Surfaces and Interfaces for Biomaterials

Surfaces and interfaces for biomaterials

Related titles from Woodhead Publishing Limited: Medical textiles and biomaterials for healthcare (ISBN 1 85573 683 7) Medical textiles and biomaterials are a significant and increasingly important part of the technical textiles industry. They cover a huge range of applications from diapers and surgical gowns to substrates for electronic sensoring of vital life signs, external use as wound care, and internal use as implantables for biodegradable post-operative support systems, as well as the replacement of body parts through tissue engineering by supplying the structure for the growth of new cells. Even the humble plaster has the potential to deliver a powerful healthcare effect through its specific skincare characteristics and controlled delivery of medications. Medical textiles and biomaterials for healthcare will discuss recent developments in the main aspects of healthcare and medical textiles covering materials, manufacture, performance, applications, standards and user experience. It will serve as an essential resource for healthcare and medical textiles manufacturers, clinicians, researchers and consumers. Biomaterials, artificial organs and tissue engineering (ISBN 185573 737 X) Maintaining quality of life in an ageing population is one of the great challenges of the twenty-first century. This book and collection of illustrated CD lectures summarises how this challenge is being met by multi-disciplinary developments in specialty biomaterials, devices, artificial organs and in-vitro growth of human cells as tissue engineered constructs. Part A provides an introduction to living and man-made materials for the nonspecialist; Part B is an overview of clinical applications of various biomaterials and devices; Part C summarises the bioengineering principles, materials and designs used in artificial organs; Part D presents the concepts, cell techniques, scaffold materials and applications of tissue engineering; and Part E provides an overview of the complex socioeconomic factors involved in technology-based healthcare, including regulatory controls, technology transfer processes and ethical issues. Each chapter is supplemented with illustrated lectures and study questions in an easy to use CD to aid the reader in self-paced instruction. Hyaluronan (two-volume set: ISBN 1 85573 570 9) (Proceedings of an international meeting, September 2000, North East Wales Institute, UK.) Hyaluronan and its derivatives have developed very quickly in the last few years from a scientific novelty into an important new material for a diverse range of medical and biomaterial applications. The landmark conference on which this two-volume reference work is based focused on recent developments and applications in the use of hyaluronan in tissue repair and reconstruction, drug delivery systems, anti-cancer treatments and joint recovery and engineering. The entire range of hyaluronan progress is dealt with in depth by more than 135 individual papers presented in two volumes covering: analytical chemistry; chemical modification; physical characterisation; cell biology and medical applications. Details of these books and other Woodhead Publishing titles can be obtained by: · visiting our website at www.woodheadpublishing.com · contacting Customer Services (e-mail: [email protected]; fax: +44 (0) 1223 893694; tel.: +44 (0) 1223 891358 ext. 30; address: Woodhead Publishing Limited, Abington Hall, Abington, Cambridge CB1 6AH, England)

Surfaces and interfaces for biomaterials Edited by Pankaj Vadgama

Published by Woodhead Publishing Limited Abington Hall, Abington Cambridge CB1 6AH England www.woodheadpublishing.com Published in North America by CRC Press LLC 2000 Corporate Blvd, NW Boca Raton FL 33431 USA First published 2005, Woodhead Publishing Limited and CRC Press LLC ß 2005, Woodhead Publishing Limited The authors have asserted their moral rights. Every effort has been made to trace and acknowledge ownership of copyright. The publishers will be glad to hear from the copyright holders whom it has not been possible to contact concerning the following: Figs 1.11 and 1.13. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from the publishers. The consent of Woodhead Publishing Limited and CRC Press LLC does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited or CRC Press LLC for such copying. Trademark notice: product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress Woodhead Publishing Limited ISBN-13: 978-1-85573-930-7 Woodhead Publishing Limited ISBN-10: 1-85573-930-5 CRC Press ISBN 0-8493-3446-6 CRC Press order number: WP3446 Project managed by Macfarlane Production Services, Markyate, Hertfordshire ([email protected]) Typeset by Godiva Publishing Services Ltd, Coventry, West Midlands Printed by TJ International Limited, Padstow, Cornwall, England

Contents

Contributor contact details

xv

Preface

xxi

Part I Forming methods 1

Fundamental properties of surfaces

P W E I G H T M A N and D S M A R T I N , The University of

3

Liverpool, UK

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 1.10 1.11

Introduction Experimental considerations Surface characteristics Active sites and kinetics Controlling crystal growth: semiconductor technology Heterogeneous catalysis Real surfaces: theoretical advances Real surfaces: experimental approaches Insight into the biological activity of surfaces Conclusion References

3 4 14 15 17 20 23 24 25 27 27

2

Control of polymeric biomaterial surfaces

29

2.1 2.2 2.3 2.4 2.5 2.6

Introduction Preparation of polymers The solid state and structure Polymer-solvent interactions The polymeric surface and surface-bulk difference The general properties of a biomaterial surface

29 29 32 35 38 39

V H A S I R C I and N H A S I R C I , METU, Turkey

vi

Contents

2.7 2.8 2.9 2.10 2.11

Modification of polymer surfaces Surface analysis Surface properties and biomaterials applications Conclusion References

3

Organic thin film architectures: fabrication and properties M C P E T T Y , University of Durham, UK

3.1 3.2 3.3 3.4 3.5 3.6 3.7

4

4.1 4.2 4.3 4.4 4.5 4.6 4.7 4.8

5

5.1 5.2 5.3 5.4 5.5 5.6 5.7

Introduction Established deposition methods Molecular architectures Molecular organization in thin films Future trends Further information References

Membranes and permeable films

N A H O E N I C H , University of Newcastle upon Tyne, UK and D M A L I K , Loughborough University, UK

40 50 56 57 57

60 60 61 65 71 78 79 79

83

Introduction 83 Materials and applications 83 Membrane characterisation 89 Blood material contact 92 Biological events at the membrane and thin film blood interface 93 Improvement of biocompatibility 99 Conclusion 99 References 100

Stable use of biosensors at the sample interface

103

J F G A R G I U L I , University of London, UK, A G I L L and G L I L L I E , University of Manchester, UK, M S C H O E N L E B E R , University of London, UK, J P E A R S O N , University of Manchester, UK, G K Y R I A K O U and P V A D G A M A , University of London, UK

Introduction Biosensor limitations Biocompatibility Materials interfacing strategy Membrane systems used in biosensors Microflows as surrogate, renewable barrier films Microfluidics and biosensors

103 104 105 116 119 138 140

Contents 5.8 5.9 5.10

6

Conclusion Acknowledgements References

Micro- and nanoscale surface patterning techniques for localising biomolecules and cells: the essence of nanobiotechnology

Z A D E M O V I C and P K I N G S H O T T , The Danish Poymer

vii 146 147 147

150

Centre, Denmark

6.1 6.2 6.3 6.4 6.5 6.6 6.7

Introduction Lithographic patterning with photons, particles and scanning probes Soft lithographic techniques Colloidal-based fabrication techniques Template-imprinted nanostructured surfaces Conclusion References

150 152 158 168 169 169 171

Part II Measurement, monitoring and characterisation 7

Surface spectroscopies

M M E H L M A N N and G G A U G L I T Z , University of

183

Tuebingen, Germany

7.1 7.2 7.3 7.4 7.5 7.6

8

Introduction Surfaces Optical detection methods Biomolecular interaction analysis Conclusion References

Surface microscopies

C Z I E G L E R , University of Kaiserslautern, and Institut fuÈr

183 183 186 190 196 197

200

OberflaÈchen- und Schichtanalytik GmbH, Kaiserslautern, Germany

8.1 8.2 8.3 8.4 8.5 8.6 8.7

Introduction Electron microscopies Scanning probe microscopies Optical microscopies Future trends Further information References

200 208 211 218 219 220 220

viii

9

Contents

Nanoindentation

A B M A N N , The State University of New Jersey, USA

9.1 9.2 9.3 9.4 9.5 9.6 9.7 9.8 9.9

Introduction Instrumentation Data analysis Thin films and coatings Hard biological materials Soft biological materials Conclusion Further information References

10

Surface plasmon resonance

10.1 10.2 10.3 10.4 10.5 10.6 10.7

Introduction Surface plasmon resonance phenomenon Surface functionalization Applications Conclusion Acknowledgements References

11

Ellipsometry for optical surface study applications

V H P EÂ R E Z - L U N A , Illinois Institute of Technology, USA

Y M G E B R E M I C H A E L and K T V G R A T T A N , City

225 225 225 229 235 236 240 241 242 242

248 248 249 257 261 264 265 265

271

University, London

11.1 11.2 11.3 11.4 11.5 11.6 11.7

12 12.1 12.2 12.3 12.4 12.5 12.6

Introduction History of ellipsometry and polarisation control Fibre based polarisation modulated ellipsometry A high birefringence fibre polarisation modulation ellipsometry Future trends Sources of further information References

Neutron reflection

J R L U , UMIST, UK

Introduction Neutron reflection and deuterium labelling Peptide interfacial assembly Lysozyme adsorption: the effect of surface chemistry Effect of size of globular proteins on their adsorption Conclusion

271 278 282 285 292 292 294

299 299 300 303 305 314 315

Contents 12.7 12.8

13

ix

Acknowledgements References

317 317

Microgravimetry

322

S V M I K H A L O V S K Y , University of Brighton, UK, V M G U N ' K O , Institute of Surface Chemistry, Ukraine, K D P A V E Y , University of Brighton, UK, P E T O M L I N S , National Physical Laboratory, UK and S L J A M E S , University of Brighton, UK

13.1 13.2 13.3 13.4 13.5 13.6 13.7 13.8 13.9 13.10 13.11 13.12

Introduction Quartz crystal microbalance technique Analytical applications of QCM Combination of QCM with other techniques Acoustic/piezoelectric sensors Future development of piezoelectric sensors Thermal gravimetry Non-QCM adsorption methods Dynamic contact angle measurements Conclusion Acknowledgements References

322 322 332 355 357 359 360 361 362 366 366 366

Part III Surface interaction and in-vitro studies 14

Interaction between biomaterials and cell tissues

Y I W A S A K I and N N A K A B A Y A S H I , Tokyo Medical

389

and Dental University, Japan

14.1 14.2 14.3 14.4 14.5 14.6 14.7

15 15.1

Introduction Surface properties of biomedical materials Surface analyses of biomedical materials Design for non-biofouling surface How to connect tissues with biomaterials Conclusion References

389 389 393 399 405 410 411

Blood flow dynamics and surface interactions

414

Clinical application and problems of medical devices in contact with blood

414

W V A N O E V E R E N , University of Groningen, The Netherlands

x

Contents

15.2 15.3

Surface interactions of blood Role of blood cells during flow: rolling of cells, effect of concentration of erythrocytes, expression of adhesive cell receptors Biomaterial surface characteristics in relation to haemocompatibility and clinical applications Haemocompatibility of metals, ceramics and polymers Biological surface treatment to improve haemocompatibility ISO 10993 requirements for testing of medical devices: simulation of clinical application including flow, blood composition, anticoagulants Test models: static, low flow, arterial flow, pulsatile/laminar flow Conclusion References

15.4 15.5 15.6 15.7 15.8 15.9 15.10

16

Cell guidance through surface cues

A K V O G T - E I S E L E , Max-Planck Institute for Polymer Research, Germany, A O F F E N H AÈ U S S E R , Institute for Thin Films and Interfaces, Research Centre JuÈlich, Germany and W K N O L L ,

418 422 423 425 427 431 432 434 435

447

Max-Planck Institute for Polymer Research, Germany

16.1 16.2 16.3 16.4 16.5 16.6

17 17.1 17.2 17.3 17.4 17.5 17.6 17.7

Introduction Surface functionalization Patterning of chemical surface cues Synaptic connections in patterned neuronal networks: communication along predefined pathways Conclusion References

Controlled cell deposition techniques

C M A S O N , University College London, UK

Introduction In-vivo and in-vitro cell interactions Two-dimensional controlled cell deposition techniques Three-dimensional controlled cell deposition techniques Future trends Further information References

447 451 454 458 461 462

465 465 466 468 480 483 484 486

Contents

18 18.1 18.2 18.3 18.4 18.5 18.6 18.7

Biofouling in membrane separation systems

Z C U I and Y W A N , University of Oxford, UK

Introduction Membrane separation ± concepts and applications Fouling mechanisms and factors affecting fouling Biofouling Fouling control Conclusion and future trends References

xi

493 493 495 500 508 513 529 530

Part IV Surface interactions and in-vivo studies 19

Bioactive 3D scaffolds in regenerative medicine: the role of interface interactions

J R J O N E S and L L H E N C H , Imperial College London, UK 19.1 19.2 19.3 19.4 19.5 19.6 19.7 19.8 19.9 19.10

Introduction The need for biomedical materials and implants Surgical procedures for bone repair Surgical procedures in lung repair A new direction: regenerative medicine Bone regeneration Tissue engineering of the lung Conclusion Acknowledgements References

20

Intravascular drug delivery systems and devices: interactions at biointerface

K S R A O , Nebraska Medical Center, USA, A K P A N D A ,

545 545 545 547 550 551 552 561 567 568 568

573

National Institute of Immunology, India and

V L A B H A S E T W A R , Nebraska Medical Center, USA 20.1 20.2 20.3 20.4 20.5 20.6 20.7 20.8

Introduction Biomaterials and biointerface Intravascular drug delivery systems Nanoparticles as an intravascular delivery system Stents Vascular grafts and catheters Future trends References

573 573 575 575 580 581 581 582

xii

Contents

21

Surface degradation and microenvironmental outcomes C C C H U , Cornell University, USA

21.1 21.2 21.3 21.4

585 585 587 589

21.5 21.6

Introduction Chemistry of synthetic biodegradable biomaterials In-vitro degradation of synthetic biodegradable biomaterials In-vivo biodegradation of synthetic biodegradable biomaterials and cell/biomaterial surface interaction Conclusion References

22

Microbial biofilms and clinical implants

619

22.1 22.2

Introduction Epidemology and costs of infection associated with clinical implants Microbiology of clinical implant infections Molecular mechanisms underlying biofilm formation Determinants of biofilm antibiotic resistance Consequences of biofilm formation on clinical implants Clinical implant infection Prevention of biofilm formation on clinical implants Further research Information resources References

619

22.3 22.4 22.5 22.6 22.7 22.8 22.9 22.10 22.11

23

601 614 614

M M I L L A R , Barts and the London School of Medicine and Dentistry

Extracellular matrix molecules in vascular tissue engineering C M K I E L T Y , D V B A X , N H O D S O N and M J S H E R R A T T , Wellcome Trust Centre for Cell-Matrix

620 621 622 624 624 625 627 630 630 631

637

Research, UK

23.1 23.2 23.3 23.4 23.5 23.6 23.7 23.8 23.9

Introduction Natural blood vessels Vascular tissue engineering Coating ECM molecules on surfaces ± a cautionary tale Biological seeding materials ECM-regulated delivery of therapeutic growth factors Future trends Acknowledgements References

637 638 641 646 651 657 658 658 659

Contents

24

Biomineralisation processes

24.1 24.2

Introduction `Biologically-induced' and `organic matrix-mediated' mineralisation Organic macromolecules Control over crystal structure Control over crystal orientation Control over morphology Control over mechanical properties Conclusion Further information References

24.3 24.4 24.5 24.6 24.7 24.8 24.9 24.10

25

F C M E L D R U M , University of Bristol, UK

xiii

666 666 667 667 671 674 676 686 687 688 688

On the topographical characterisation of biomaterial surfaces 693 P E T O M L I N S and R L E A C H , National Physical Laboratory, UK, P V A D G A M A , University of London, UK, S M I K H A L O V S K Y and S J A M E S , University of Brighton, UK

25.1 25.2 25.3 25.4 25.5 25.6 25.7 25.8 25.9 25.10 25.11 25.12 25.13 25.14

Introduction Biomaterials, surfaces and biocompatibility What is a surface? Surface measurement Filters Quantifying surface texture Two-dimensional profile data Three-dimensional data Techniques for surface texture measurement Traceability and calibration Conclusion Further reading Acknowledgements References

693 693 694 694 695 696 696 700 703 713 713 714 715 715

Part V Appendices 26

Surface modification of polymers to enhance biocompatibility

719

Introduction Polymers in medical applications

719 720

M T A V A K O L I , TWI Limited, UK

26.1 26.2

xiv

Contents

26.3 26.4 26.5 26.6 26.7

Biocompatibility Surface modification techniques Future trends Acknowledgements References

27

Issues concerning the use of assays of cell adhesion to biomaterials

S L J A M E S and S M I K H A L O V S K Y , University of Brighton, UK, P V A D G A M A , University of London, UK and P E T O M L I N S , National Physical Laboratory, UK

722 723 740 741 741

745

27.1 27.2 27.3 27.4 27.5 27.6 27.7

Introduction Measurement objectives Issues of interpretation of adhesion measurements Sources of variability in adhesion assays Methods of assaying cell adhesion in current use Conclusion References

28

Protein adsorption to surfaces and interfaces

28.1 28.2 28.3 28.4 28.5 28.6 28.7 28.8

Introduction Classification of biomaterials surfaces and interfaces Non-specific adsorption to hard surfaces General rules of non-specific adsorption to flat surfaces Non-specific adsorption to `soft' surfaces Non-specific adsorption to penetrable surfaces and interfaces Future trends References

763 764 765 768 773 774 775 776

Index

782

B M I L T H O R P E , University of New South Wales, Australia

745 746 750 753 757 760 761

763

Contributor contact details

Introduction Professor Pankaj Vadgama Director, IRC in Biomedical Materials Queen Mary University of London Mile End Road London E1 4NS UK E-mail: [email protected] Chapter 1 Professor Peter Weightman and Dr David S Martin Physics Department and Surface Science Research Centre The University of Liverpool Liverpool L69 3BX UK Tel: +44 (0)151 794 3871 E-mail: [email protected] E-mail: [email protected] Chapter 2 Professor V Hasirci Department of Biological Sciences Biotechnology Research Unit and Professor N Hasirci

Department of Chemistry METU Ankara 06531 Turkey E-mail: [email protected] E-mail: [email protected] Chapter 3 Professor M C Petty School of Engineering and Centre for Molecular and Nanoscale Electronics University of Durham South Road Durham DH1 3LE UK Tel: +44 (0)191 334 2419 E-mail: [email protected] Chapter 4 Dr Nicholas A Hoenich School of Clinical Medical Sciences Medical School University of Newcastle Framlington Place Newcastle upon Tyne NE2 4HH UK Tel: +44 (0)191 222 6998 E-mail: [email protected]

xvi

Contributor contact details

Dr Danish J Malik Department of Chemical Engineering Loughborough University Loughborough LE11 3TU UK Tel: +44 (0)1509 222507 E-mail: [email protected] Chapter 5 Joseph Gargiuli IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 5547/3254. E-mail: [email protected] Andrew Gill Inverness Medical Ltd Beechwood Business Park North Inverness IV2 3ED UK Tel: +44 (0)1463 721706. E-mail: [email protected] Monika Schoenleber IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 3316 E-mail: [email protected] J Pearson Department of Epidemiology University of Manchester Manchester M13 9PL UK

G Kyriakou IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 3316 E-mail: [email protected] Professor Pankaj Vadgama Director, IRC in Biomedical Materials Queen Mary, University of London Mile End Road London E1 4NS UK Tel: +44 (0)20 7882 3316 E-mail: [email protected] Chapter 6 Dr P Kingshott The Danish Polymer Centre Riso National Laboratory PO Box 49 Fredericksborgvej 399 DK-4000 Roskilde Denmark E-mail: [email protected] E-mail: [email protected] Chapter 7 Professor G Gauglitz and M Mehlmann Institut fuÈr Physikalische und Theoretische Chemie UniversitaÈt TuÈbingen Auf der Morgenstelle 9 72076 TuÈbingen Germany

Contributor contact details E-mail: [email protected] E-mail: [email protected] Chapter 8 C Ziegler Department of Physics, University of Kaiserslautern and Institut fuÈr OberflaÈchen- und Schichtanalytik GmbH Erwin-SchroÈdinger-Straûe 56 67663 Kaiserslautern Germany Tel: +49 631 205 2855 E-mail: [email protected] Chapter 9 Adrian Mann Department of Biomedical Engineering Rutgers The State University of New Jersey Piscataway NJ 08854 USA E-mail: [email protected] Chapter 10 Professor Victor Perez Luna Room 144 Perlstein Hall Illinois Institute of Technology in Bioengineering, Chicago USA E-mail: [email protected]

xvii

Chapter 11 Dr Y M Gebremichael and Professor K T V Grattan School of Engineering and Mathematical Sciences City University Northampton Square London EC1V 0HB UK Tel: +44 (0)20 7040 3888 Fax: +44 (0)20 7040 8568 E-mail: [email protected] E-mail: [email protected] Chapter 12 Professor Lu Biological Physics Group School of Physics and Astronomy The University of Manchester G13, Sackville Street Building Sackville Street Manchester M60 1QD UK Tel: +44 (0)161 2003926 E-mail: [email protected] Chapter 13 Professor S V Mikhalovsky, Dr K D Pavey and Dr S L James School of Pharmacy and Biomolecular Sciences, University of Brighton Cockcroft Building Lewes Road Brighton BN2 4GJ UK Tel: +44 (0)1273 642034 Fax: +44 (0)1273 642115 E-mail: [email protected]

xviii

Contributor contact details

Tel: +44 (0)1273 642042 Fax: +44 (0)1273 679333 E-mail: [email protected] Dr V M Gun'ko Institute of Surface Chemistry 17 General Naumov Street 03164 Kiev Ukraine Tel/Fax: +38 (044) 5123095 E-mail: [email protected] Dr P Tomlins Materials Centre, National Physical Laboratory, Queens Road, Teddington, Middlesex, TW11 0LW UK Tel: +44 (0)208 943 6778 Fax: +44 (0)208 943 6453 E-mail: [email protected] Chapter 14 Professor N Nakabayashi Institute of Biomaterials and Bioengineering Tokyo Medical and Dental University 2-3-10 Kanda-surugadai Chiyoda-ku Tokyo 101-0062 Japan Tel: +81-47-341-9734 E-mail:[email protected] Dr Yasuhiko Iwasaki Institute of Biomaterials and Bioengineering Tokyo Medical and Dental University

Kanda-surugadai Chiyoda-ku Tokyo 101-0026 Japan Tel: +81-3-5280-8026 E-mail: [email protected] Chapter 15 Dr Willem van Oeveren, Dept of Biomedical Engineering, University Medical Center Groningen, Ant. Deusinglaan 1 PO Box 196, 9700AD Groningen The Netherlands Tel: +31 503633127 E-mail: [email protected] Chapter 16 Professor W Knoll Max-Planck-Institut fuÈr Polymerforschung Postfach 3148 55021 Mainz Germany Tel: +49 06131 / 379 160 E-mail: [email protected] Dr Angela K Vogt-Eisele Cell Physiology ND/4 Ruhr University Bochum Universitystr. 150 44 780 Bochum Germany Tel: +49 (0)234 32 26756 Fax: +49 (0)234 32 14129 E-mail: [email protected]

Contributor contact details Professor Andreas OffenhaÈusser Institute for Thin Films and Interfaces (ISG-2) Forschungszentrum JuÈlich D-52425 JuÈlich Germany Tel: +49 (0)2461 612330 Fax: +49 (0)2461 618733 E-mail: [email protected] Chapter 17 Mr Chris Mason FRCS Department of Biochemical Engineering University College London Torrington Place London WC1E 7JE UK Tel: +44 (0)20 7679 0140 Fax: +44 (0)20 7209 0703 E-mail: [email protected] Chapter 18 Professor Zhanfeng Cui Department of Engineering Science Oxford University Parks Road Oxford OX1 3PJ UK Tel: +44 (0)1865 273118 E-mail: [email protected] Dr Yinhua Wan Department of Engineering Science Oxford University Parks Road Oxford OX1 3PJ UK Tel: +44 (0)1865 273059 E-mail: [email protected]

xix

Chapter 19 Professor Larry Hench and Dr Julian Jones Department of Materials Imperial College London South Kensington Campus London SW7 2AZ UK E-mail: [email protected] E-mail: [email protected] Chapter 20 Vinod Labhasetwar, Ph.D. College of Pharmacy Department of Pharmaceutical Sciences 986025 Nebraska Medical Center Omaha, NE 68198-6025 USA Tel: (402) 559-9021 E-mail: [email protected] Chapter 21 C Chu Department of Textiles and Apparel, and Biomedical Engineering Program Martha Van Rensselaer Hall Cornell University Ithaca, New York, 14853-4401 USA Tel: 607-255-1938 E-mail: [email protected] Chapter 22 Michael Millar Department of Microbiology Barts and the London School of Medicine and Dentistry Whitechapel

xx

Contributor contact details

London E1 2AD UK Tel: +44 (0)20 7377 7080 E-mail: [email protected] Chapter 23 Cay M Kielty, Daniel V Bax, Nigel Hodson and Michael J Sherratt School of Biological Sciences University of Manchester Michael Smith Building Oxford Road Manchester M13 9PT UK E-mail: [email protected] Chapter 24 Dr Fiona Meldrum School of Chemistry University of Bristol Cantock's Close Bristol BS8 1TS UK Tel: +44 (0)117 3317215 E-mail: [email protected] Chapter 25 Dr Paul Tomlins National Physical Laboratory Queens Road Teddington Middlesex TW11 0LW UK E-mail: [email protected] Professor Pankaj Vadgama Director, IRC in Biomedical Materials Queen Mary

University of London Mile End Road London E1 4NS UK E-mail: [email protected] Chapter 26 Professor Mehdi Tavakoli TWI Limited Granta Park Abington Cambridgeshire CB1 6AL UK E-mail: [email protected] Chapter 27 Dr S L James School of Pharmacy and Biomolecular Sciences University of Brighton Cockcroft Building Lewes Road Brighton BN2 4GJ UK Tel: +44 (0)1273 642042 Fax: +44 (0)1273 679333 E-mail: [email protected] Chapter 28 Professor B Milthorpe Head of School Graduate Scool of Biomedical Engineering University of New South Wales UNSW Sydney 2052 Australia Tel: +61-2-9385-3911 Fax: +61-2-9663-2108 E-mail: [email protected]

Preface

Biomaterials research has undergone a variety of evolutionary, one might say revolutionary, developments in recent years. The full entry of biomaterials as a credible modality in modern medicine was essentially triggered by the work of John Charnley on hip replacement. Then, as now, bulk materials properties and biomechanics took centre stage in view of the stringent mechanical and tribological demands of the implants. However, such issues cannot be the sole determinants of clinical outcome. Interest in bulk properties has inevitably shifted to the important consideration of the surface with the plethora of allied interfacial phenomena, conditioning clinical outcome. A consideration of the interface goes beyond the realms of an intellectual research exercise given that, even today, implants are better characterised by a bioincompatibility, rather than the hoped for `inertness' of biocompatibility. In mass terms, of course the surface is insignificant compared with the bulk. However, whilst the latter is a largely sequestered environment, amenable to at least some degree of engineering/physical science based prediction, the interface is the seat of some extraordinarily powerful, frontline biological responses. These variously conspire to attack the `non-self' implant firstly to try to degrade and eliminate it, and if that is not possible, to mask it off with an inert capsule; not a very promising outcome for precision engineering based therapeutics. The rules that govern the complex biological events are yet to be fully worked out, but it is a fair assumption that study of surface chemistry, organisation, topography, mechanics and physics, etc., will repay the effort. Unfortunately, the importance of surface processes in determining functional outcomes is not yet matched by the attention given in dedicated texts. The growing body of original literature also seems increasingly dissipated in analytical, chemical and biophysical journals. This book attempts to redress the information imbalance between surface and bulk descriptions or biomaterials. The series of chapters attempts to cover a spectrum from the fundamentals of surface structure and forming methods to biological and clinical outcomes. The link between experimental studies of surfaces in a controlled biomatrix and clinical results is a difficult one to make,

xxii

Preface

but at least the authoritative descriptions provided here will promote fresh thinking in this important area. Without measurement tools, understanding is restricted, so some key surface analytical techniques, both classical and developing (viz. probe microscopies, optics) are included. Ultimately, quantitative measurement is vital to understanding and so this book is in part an expression of the UK interest in metrology, represented through Department of Trade and Industry (DTI) programmes on materials metrology. I would like to express my sincere thanks to the many authors who were so patient with me and who contributed so effectively in making this book a viable exercise. My thanks also to Francis Dodds, Melanie Cotterell and their colleagues at Woodhead, whose continuous support and advice made the whole process both pleasurable and interesting and one that was so ably initiated by Gwen Jones. Finally my thanks to Catherine Jones of the IRC in Biomedical Materials for her unfailing assistance over the innumerable author interactions in modifying and adapting texts and for ensuring the project remained on track. Pankaj Vadgama IRC in Biomedical Materials University of London

Part I

Forming methods

1

Fundamental properties of surfaces P W E I G H T M A N and D S M A R T I N , The University of Liverpool, UK

1.1

Introduction

Surface scientists occasionally observe that while God created solids, the Devil created surfaces. This phrase encapsulates the fact that while surfaces often dominate the behaviour of materials they are very difficult to study. Surfaces usually lack the high symmetry and purity of the interior of a solid and are often strongly influenced by adsorbed impurities from their environment. It is hard to overstate the significance of the Devil's work since surfaces are very important. Phenomena experienced in everyday life such as corrosion, adhesion, adsorption, friction and lubrication all occur at surfaces. More intimately, the crucial role played by surfaces in biocompatibility gives them an importance in the design of materials used in dentistry, contact lenses and medical implants such as hip joints and knee replacements. Industrial processes that occur at surfaces have a great impact on our lives and include crystal growth, semiconductor device manufacture and heterogeneous catalysis. Surface properties will also have a dominant influence on the emerging field of nanotechnology. The control of surface properties is thus essential to the function of a wide variety of materials. Clearly, then, a primary aim of surface scientists is to obtain a sufficient understanding of surfaces to make it possible to control surface properties. The field of surface science concerns fundamental, nanoscale investigations of surface phenomena that are both scientifically important and technologically relevant. The subject is relatively new since the experimental study of clean surfaces was delayed for many years by the crucial limitation that in a pressure in excess of 10ÿ6 mbar a clean surface adsorbs a monolayer of impurities within a few seconds. Thus experimental work on clean surfaces could not begin until the development in the 1960s of ultra-high vacuum (UHV) techniques, which reach pressures of 10ÿ10 mbar or less. This technological advance led to a rapid increase in surface studies and a proliferation of experimental techniques of which there is space here to give only the briefest of descriptions of some of the most important.

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The advances made in the understanding of surfaces in UHV conditions in the last forty years raise two questions. Firstly, to what extent are the surfaces that we now understand representative of the field in general? Secondly, does an understanding of the behaviour of surfaces in UHV provide a good guide to the behaviour of surfaces in the ambient conditions in which we find them in everyday life? The second of these questions is particularly pertinent to the role of surfaces in biomaterial applications and it will be addressed later in this chapter. The first question requires a rough estimate of how many surfaces are worthy of study since there are an infinite number of ways of terminating single crystals, to say nothing of the number of possible surfaces of amorphous materials. As a rough guide we note that Wyckoff's six volume classification1 lists ~7000 crystal structures. If we assume that each crystal structure has three important crystal faces and that it is appropriate to seek to understand the interactions that occur between each of these faces with the ten most important gases then we have a target for surface science of understanding ~200,000 surfaces. So far, the structures of ~1000 surfaces have been determined, about 1% of the target. We have a good understanding of the behaviour of the known surface structures in UHV and some understanding of the factors that are important in catalysis. In addition surface science has made major contributions to the considerable progress that has been made in the design and controlled growth of semiconductor systems. However, as will be made clear later, we are only just beginning to develop an understanding of surfaces in ambient conditions. The study of surfaces thus has a long way to go, particularly in addressing issues that are important in the real world.

1.2

Experimental considerations

We begin with a brief account of the importance of UHV and a description of single crystal surfaces. There are many experimental probes capable of detailed investigations of surfaces and interfaces, however, we have space to give only a brief overview of some of the most commonly used surface techniques. A more extensive introduction to these and many other techniques may be found in the books by Woodruff and Delchar,2 Zangwill,3 Prutton4 and Venables.5

1.2.1 Ultra-high vacuum As indicated in the introduction, the experimental investigation of the fundamental properties of surfaces have mostly taken place in UHV where the pressure is typically 10-10 mbar ± thirteen orders of magnitude lower than atmospheric pressure. A UHV environment is required to prepare a well-defined clean surface and maintain it for a sufficient time for experimental studies. In addition to surface preparation, a good vacuum is also a prerequisite for many of the experimental probes used to study surfaces since these probes are often

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5

based upon controlling the trajectories of electrons and ions. Vacuum technology has developed pumping systems capable of maintaining a UHV environment within stainless steel chambers for indefinite periods of time. Experiments are not exclusively performed on clean surfaces in UHV and the recent development of experimental probes that are capable of operating in a non-UHV environment is giving rise to an increasing trend of experimental studies of surfaces in ambient and liquid environments.

1.2.2 Crystal surfaces and surface preparation The majority of experimental surface studies have been performed on single crystals in order to simplify the atomic and electronic structure of the surface. Crystal surfaces can be prepared so as to consist of relatively large flat terraces made up of atoms of similar atomic coordination, with relatively few atoms associated with defect sites such as steps. The majority of single crystals grow in one of four basic structures: simple cubic (SC), face centred cubic (FCC), body centred cubic (BCC) or hexagonal close packed (HCP). When a single crystal is terminated by a surface, then, depending on the angle of the termination, different atomic arrangements are exposed. These different surfaces are described by the Miller indices and an introduction to this system of classification of crystal structures and surfaces can be found in ref. 6, which also explains the Wood notation that is used to describe the symmetry of surfaces. For FCC and BCC crystals, surface planes are defined by three integers. The three `low index' surfaces, (110), (111) and (100) created from the FCC and BCC structures are shown in Figs 1.1 and 1.2, respectively. The figures show that different crystal planes have different atomic densities and hence differences in free energy at the surface. The free energy of a surface is an important determinant of its behaviour as will be discussed later in considerations of crystal growth. Free energy consideration also influences the natural cleavage planes of crystalline materials and cleavage along a nonpreferred direction often results in a rough morphology composed of small areas of energetically preferred faces known as faceting. The cleavage of single crystals in UHV is one of the easiest ways of producing a clean surface. However, it can only be applied to the natural cleavage planes and these are not always the most important surfaces of a material. A variety of ways have been developed for producing clean surfaces on crystal faces that cannot be obtained by cleaving. These often involve bombarding the surface with argon ions in UHV to remove impurities followed by annealing to remove the structural damage. However, there is no cleaning procedure that works for all surfaces, and a considerable amount of effort is devoted to finding a recipe for the cleaning of a surface which is to be studied for the first time. Fundamental investigations of systems in ambient air are naturally limited to surfaces that can be prepared in air, and this draws attention to structural

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1.1 Surface structures created from cleavage of the FCC structure: (a) (110), (b) (111), and (c) (100). For the surface structures, the unit cells are shown: light spheres ˆ 1st layer atoms, dark spheres ˆ 2nd layer atoms.

stability in the presence of oxygen. Surfaces like graphite and mica that are stable against oxidation and which are easily cleaved in air to yield clean surfaces have naturally attracted attention. Cleaving graphite ± a layered material ± produces a well-defined surface consisting of large atomically flat terrace planes bounded by steps (Fig. 1.3). Such a relatively simple structure is ideal for investigating adsorption of organic molecules, and graphite has been used extensively for such studies.7,8 Gold surfaces are relatively inert to the effects of the atmosphere and this element provides some of the few single

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1.2 Surface structures created from cleavage of the BCC structure: (a) (110), (b) (111), and (c) (100). For the surface structures, the unit cells are shown: light spheres ˆ 1st layer atoms, dark spheres ˆ 2nd layer atoms, darkest spheres ˆ 3rd layer atoms.

crystal surfaces that can be prepared by the flame annealing technique for use as electrodes in electrochemical experiments at the solid/liquid interface. The surface structures shown in Figs 1.1 and 1.2 represent perfect terminations of the bulk crystal structure. The periodicity of the surface structures facilitates their representation by two-dimensional unit cells, which

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1.3 (a) Atomic structure of graphite showing layers (b) AFM image of the graphite surface showing atomically flat terraces and steps.

can be repeated to represent the extended surface structure. For a perfectly terminated crystal surface, the lengths of the surface unit cell are determined by the lengths of the unit vectors of the corresponding two-dimensional plane of the three-dimensional unit cell of the bulk crystal structure. This relationship is termed a (1  1) surface structure in the Wood terminology.6 Atoms at newly created surfaces sometimes adopt different positions from those expected from a perfect termination of the bulk crystal structure and such `reconstructed' surfaces may adopt a new symmetry with a periodicity that differs from that of the bulk crystal structure. This new periodicity can often be represented relative

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1.4 The FCC(110)-(1  2) reconstruction. Light spheres ˆ 1st layer atoms, dark spheres ˆ 2nd layer atoms, darkest spheres ˆ 3rd layer atoms.

to the symmetry of the unit cell of the bulk structure by a change in the ratio of the surface to bulk unit vectors. The surface periodicity is then described in terms of these ratios, n and m, as (n  m) in the Wood notation. Sometimes the representation of surface structures also requires a rotation of the axes of the surface unit cell relative to those of the bulk unit cell. Surface structures often involve a relaxation of the first few atomic planes in the direction normal to the surface. One example of a common reconstruction is the (1  2) `missing row' structure of the FCC(110) surface. In this reconstruction, every other ‰1 1 0Š row of atoms from the normal (1  1) surface is missing (Fig. 1.4) and this doubling of periodicity in one dimension is captured by the (1  2) Wood terminology. The clean (110) surfaces of Au, Pt and Ir at room temperature in UHV all adopt the (1  2) structure. There are many different surface reconstructions and surfaces often reconstruct as a result of the adsorption of foreign species. The surface reconstructions of the Si(100) and (111) surfaces are among the most important technologically and are shown in Figs 1.5 and 1.6. These reconstructions are driven by the reactivity of the unsatisfied covalent bonds of the Si surface atoms. The reconstruction of the Si(100) surface is easy to understand in that it arises from the pairing of adjacent Si surface atoms to form dimers. The dimers form rows and since the surface periodicity along the rows is the same as that of the bulk structure while the surface periodicity at right-angles to the dimer rows is twice that of the bulk structure this is termed a (1  2) reconstruction in the Wood notation. As Fig. 1.6 shows, the reconstruction that occurs on the Si(111) surface is very complex. It has a (7  7) periodicity relative to the bulk structure and its formation requires a considerable displacement and rearrangement of the atoms in the top three layers of the

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1.5 The Si(100) 1  2 reconstruction (courtesy Klaus Hermann, Fritz-HaberInstitut, Berlin).

Si(111) surface. The determination of the atomic positions in this structure took a considerable amount of time and required the application of all the major surface science experimental techniques. It is one of the many triumphs of the field. Once a well-defined single crystal surface is prepared, controlled experiments can then be performed such as investigating the first stages of oxidation or the interaction of the surface with molecular adsorbates. Structural defects such as adatoms, vacancies and steps, in addition to chemical impurities, can dramatically affect surface properties. It is often possible to control the introduction of these defects to surfaces and observe their impact on the system under investigation.

1.6 The Si(111) 7  7 reconstruction (a) STM image (Omicron Nanotechnology GmbH) and (b) surface structure model (reprinted from ref. 6 with permission of Oxford University Press).

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1.2.3 Determination of the elemental and chemical composition of surfaces: photoelectron and Auger electron spectroscopy Photoelectron spectroscopy involves illuminating a surface with photons of fixed energy and analysing the electrons subsequently emitted from the surface region. Atomic core hole states created by the X-ray emission of photoelectrons often decay by the Auger process which results in the emission of a second electron with an energy that is characteristic of the atomic element involved. These techniques can also be used to estimate the relative concentrations of the atomic species present in surfaces and the thickness of overlayers. Chemical bonds give rise to small changes in XPS and AES spectra that can be used to obtain insight into the chemical composition of surfaces. The electronic structure of a surface can be investigated using ultra-violet (UV) photoemission spectroscopy where the lower energy and momentum of UV photons relative to high energy X-rays allows the valence levels to be measured at high resolution. Unoccupied electronic bands can be examined using inverse photoemission spectroscopy (IPES).

1.2.4 Determination of surface structures: diffraction techniques The determination of surface structures on a macroscopic scale is obtained using diffraction techniques which, by their nature, emphasise periodicity. The more commonly used of these techniques are low-energy electron diffraction (LEED) and surface X-ray diffraction (SXRD). They give information on surface crystallography and can be used to determine the atomic structure of reconstructions and the overlayer structures of adsorbates. While LEED is restricted to a UHV environment, SXRD can be used to study surfaces at solid/ liquid interfaces.

1.2.5 Determination of the nature of molecular adsorption at surfaces: vibration spectroscopies and temperature programmed desorption Vibrational spectroscopies such as Raman spectroscopy and reflection absorption infra-red spectroscopy can provide information on molecular orientations at surfaces and on molecule-surface bonding. For example, whether a planar molecule adsorbed at a surface is flat-lying or upright can be distinguished. These techniques exploit the fact that chemical bonds absorb infra-red light at characteristic vibrational frequencies. An absorption spectrum reveals the chemical species present and shifts in frequencies from those expected from the `free' molecules in solution indicate interactions with the surface or other surface species.

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Temperature programmed desorption (TPD) is primarily used to determine the binding energies of molecules adsorbed at surfaces. In this technique a linearly increasing temperature gradient is applied to the sample and the molecular mass of desorbing species is recorded as a function of time. In this way, specific mass fragments can be identified and the strength of interactions and the activation energy for desorption can be determined since the least bound molecules will desorb first as the temperature is raised.

1.2.6 Techniques for the study of surfaces in ambient conditions: scanning probe and optical techniques Scanning probe techniques such as scanning tunnelling microscopy (STM) and atomic force microscopy (AFM) provide information on atomic and electronic structure on a local scale. They reveal the morphology of the surface and the degree and extent of surface roughness and they can achieve atomic resolution. The mechanism of image formation involves the surface of interest being brought into close contact with a sharp tip and exploiting the precise scanning motion that can be obtained in the plane of the surface by the use of piezoceramic crystals. The surface is then scanned underneath the tip whilst monitoring a certain property of interaction between the tip and surface. For STM, a tunnelling current, of the order of nA, between the tip and sample is monitored and maintained at a constant value by an electronic feedback loop. The measured current is extremely sensitive to tip-sample separation and increases as the separation decreases. Thus, to maintain a constant current, the sample ± due to its morphology ± is moved up or down perpendicular to the surface plane and this information is used to construct a three-dimensional image of the surface. In AFM, the tip is mounted at the free end of a flexible cantilever that is pivoted at the other end. A laser beam reflecting off the back of the cantilever monitors the deflection of the tip-cantilever arrangement by the morphology of the surface. A constant deflection is maintained by moving the surface up or down and an image sensitive to the height is recorded. The basic principles of the operation of STM and AFM are shown in Figs 1.7 and 1.8. Optical probes have the potential to study surfaces in ambient conditions but need to allow for the fact that only ~1% of the optical signal reflected from a solid arises from the surface. This dominance of the bulk contribution to the reflected signal is overcome in reflection anisotropy spectroscopy (RAS) and the related technique of reflection anisotropy microscopy (RAM) by measuring the difference in reflectance of normal incidence linearly polarised light between two orthogonal directions in the surface plane. When applied to a cubic crystal material, the contribution from the bulk cancels by symmetry and RAS becomes a surface-sensitive probe of optical anisotropy. RAS was developed to monitor semiconductor growth at high pressures9 and has developed into a probe of metal

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1.7 Basic schematic showing the principle of operation of an STM (courtesy Michael Schmid, Technische Universitaet, Wien).

1.8 Basic schematic showing the principle of operation of an AFM.

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surfaces10 with the potential to monitor molecular assembly at surfaces.11 Nonlinear optical techniques such as second harmonic generation (SHG) and sum frequency generation (SFG) are also surface sensitive and can be used to investigate chemical reactions at surfaces in a wide range of environments.12 After this brief introduction to the practical aspects of surface science we now review some of the most important conclusions reached from the study of surfaces.

1.3

Surface characteristics

Surfaces are usually expected to be flat, rigid and inert. However, this is rarely the case for real surfaces. At a microscopic level even the smoothest crystalline surface contains steps where atomic terraces of different height meet. Furthermore, these steps usually do not have straight edges but display kinks as they meander across a surface. The atoms at steps and kinks can have very different properties from atoms in the middle of terraces, and they often determine surface properties of technological importance such as the growth of crystals and the break up and assembly of molecules in catalytic reactions. Real surfaces are often rather flexible and are able to change their atomic structure in response to interactions with molecules.13,14 We have noted that the (110) surfaces of FCC metals often adopt a (1  2) missing row reconstruction (Fig. 1.4). This occurs because energetically the surface is finely balanced between the perfectly terminated and rather open (110) crystal plane and the (1  2) structure with its lower-energy close-packed (111) facets. Quite small changes to the surface electronic structure can induce changes in the structure of (110) surfaces of FCC crystals. In an electrochemical cell for example, the (1  1) to (1  2) phase transition can be induced on an Au(110) electrode by a change in charge density resulting from a change in the electrode potential and this change is reversible. Similar reconstructions can be induced by changing the potential on the other low index faces of Au in an electrochemical environment. In UHV the (1  2) reconstruction is the equilibrium structure of the Au(110) and Pt(110) faces though small amounts of adsorbed CO will change the Pt(110) surface back to a (1  1) structure. The opposite behaviour is found by the deposition of small amounts of alkali elements on the (110) faces of Ni, Cu, Pd and Ag which induce a (1  1) to (1  2) phase transition. The adsorption of larger molecules can induce more radical changes in surface structure. A dramatic example of this is the adsorption of hexa-tertbutyl-decacyclene (C60H66) on Cu(110).15 This molecule is approximately flat and it anchors to the Cu(110) surface through the formation of a characteristic trench in which 14 Cu atoms are dug out of the surface in a staggered arrangement involving two neighbouring close-packed rows, the free Cu atoms being added to kink sites at steps. When the molecules desorb, the surface is left with these characteristic trenches. It is as if the molecule leaves behind a footprint on the surface.

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Another remarkable example of adsorbate-induced restructuring is provided by the study by Bowker and co-workers14 on the adsorption of formic acid on a partially oxidised Cu(110) surface. The adsorption of oxygen creates a well known structure with (2  1) periodicity in which Cu atoms are incorporated into linear O-Cu-O-Cu rows orientated in [001] directions which are anchored at steps and which grow into islands of the (2  1) structure as the oxygen coverage increases. The formation of this structure is facilitated by the high mobility of Cu atoms at this surface. The adsorption of formic acid leads to a reaction with the (2  1) islands in which the Cu atoms at the end of the O-CuO-Cu units are released. These Cu atoms migrate and attach to more stable sites at the step edges between adjacent (2  1) islands. The subsequent desorption of formic acid leaves a surface in which the initial almost linear steps have been transformed into a sharp and jagged sawtooth structure.

1.4

Active sites and kinetics

The term `active site' refers to a specific atomic site on a surface where an interaction or reaction occurs. Active sites often involve atoms located at steps and kinks which have different coordination numbers from those of atoms bound in the terrace planes. The nature and influence of active sites can often be determined from TPD studies as illustrated by the following example of H2 adsorbed on Pt surfaces.13 The ideal Pt(111) surface is flat with hexagonal symmetry and one can identify three obvious adsorption sites as illustrated in Fig. 1.9; the top site where a molecule sits on top of an atom, the bridge site where a molecule is located between two atoms and the three-fold hollow site between three adjacent atoms. TPD studies of hydrogen desorption from the flat (111) surface show two overlapping peaks (Fig. 1.10) indicating that hydrogen occupies two sites on Pt(111) terraces with very similar binding energies. The Pt(557) surface consists of flat (111) terraces and steps aligned along a common direction giving a surface that is rather like a staircase (Fig. 1.9b). TPD results for this surface show peaks in the desorption of hydrogen that occur at two very different temperatures (Fig. 1.10). The peak at lower temperature corresponds to the peaks seen on the Pt(111) surface and is associated with adsorption on the (111) terraces and the peak at higher temperatures is associated with hydrogen that is more tightly bound to the higher coordination sites available at the steps. The TPD results for the desorption of hydrogen from the Pt(12,9,8) surface, which consists of (111) terraces and steps with the steps containing a large number of kink sites (Fig. 1.9c), is shown in Fig. 1.10. An additional peak is observed in the TPD results for this surface compared to the results for the Pt(557) surface (Fig. 1.10) and it is natural to associate this higher temperature peak with the binding of hydrogen to kink sites. Some insight into the nature of the adsorption of hydrogen on Pt surfaces can be obtained from TPD studies of surfaces which are dosed with a mixture of H2

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1.9 Surface structures of (a) Pt(111) showing the three possible adsorption sites, (b) Pt(557) and (c) Pt(12,9,8).

and D2. If the molecules dissociate on the surface it is possible that HD molecules will form prior to desorption. It is found that on well ordered Pt(111) surfaces the formation of HD is below the detection limits of the experiment indicating that the molecules do not dissociate on adsorption. However on stepped surfaces the production of HD indicates that all the molecules dissociate on adsorption. Clearly the symmetry and coordination of the adsorption site has an important influence on the details of the adsorption and desorption processes, a result that is borne out by a large number of studies on a wide variety of other systems.

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1.10 TPD results of H desorption from Pt(111), Pt(557) and Pt(12,9,8) (reprinted from ref. 13 with permission of John Wiley & Sons Inc.).

1.5

Controlling crystal growth: Semiconductor technology

One of the most impressive achievements of surface science has been in fostering the growth of semiconductor devices that in the last twenty years have revolutionised many aspects of technology particularly in the field of communication. This industry was not possible until the development of UHV technology and associated surface science techniques of surface preparation and characterisation. Indeed, all stages in the development of semiconductor devices have benefited from the exploitation of advances made in surface science. The key to this technology is establishing control over the growth of single crystal semiconductors and this is usually achieved by the molecular beam epitaxy (MBE) technique, which operates in a UHV environment, or by a variety of chemical vapour deposition (CVD) techniques which operate at higher pressures but in very controlled conditions. We can obtain some insight into the factors that are important in controlling crystal growth by considering the growth of crystals from solution in which the differences in the free energy of different crystal faces have a strong influence on the crystal shape. Provided there is sufficient kinetic freedom during the growth process then atoms being

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added to a crystal during growth will migrate to the face with the lowest free energy. This face will thus grow faster than other faces and, all other factors being equal, the equilibrium crystal shape will be the one in which the relative distances of each face from the nucleation centre is inversely proportional to the ratio of the free energies of the faces. Often, however, the `other factors' are not equal and crystal shape can be strongly influenced by factors such as temperature and the composition of the growth medium. This was discovered several hundred years ago when in 1783 Rome de l'Isle showed that octahedrons are formed instead of cubes if NaCl is grown in the presence of urine.16 This puzzling observation has recently been explained17 in terms of the role of the urea molecule in stabilising the (111) crystal face in solution during growth. A similar rebalancing of kinetic factors caused by subtle changes in the conditions of growth probably explains why natural diamonds tend to have octohedral shapes while those that are grown under high pressure and high-temperature conditions in the laboratory have a more cubic form though the morphology of the latter can be modified to a certain extent by varying the growth temperature. Diamond is also a good illustration of the fact that the structure of a solid is not always in equilibrium since the phase diagram of carbon (Fig. 1.11) shows that at normal pressures and temperatures the equilibrium structure of carbon is graphite, which consists of layers of strongly bonded carbon atoms (Fig. 1.3) with much weaker bonding between the layers. The cubic diamond form in which each carbon is at the centre of a tetrahedron of neighbouring carbon atoms (Fig. 1.12) is thus metastable at room temperature and pressure. The reason diamond does not readily transform into graphite is due to the energy barrier that

1.11 The phase diagram of carbon.

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1.12 The diamond crystal structure (reprinted from Introduction to Solid State Physics, 7th edition (1996) by C. Kittel with permission from John Wiley & Sons Inc.).

has to be overcome to re-orientate the inter-atomic bonds. The system is thus not able to reach equilibrium because kinetic factors prevent it from escaping from a local energy minimum. These kinetic factors can be exploited in the growth of diamond. The carbon phase diagram (Fig. 1.11) shows that it is reasonable to expect to be able to grow diamond at the high-pressure, high-temperature region on the border between the graphite and diamond phases. However, it is surprising that diamond can also be grown in the low-pressure and relatively low-temperature region by the chemical vapour deposition (CVD) process, which makes use of a microwave plasma excitation of a hydrocarbon vapour stream. The key to the CVD growth of diamond is to include a very low concentration of carbon in the feedstock gas: a typical growing medium consists of 99.5% H2 and 0.5% CH4. This ensures that the plasma is dominated by H ions and these ions terminate the surface carbon atoms of the diamond ensuring the maintenance of local tetrahedral and near tetrahedral bonding configurations. When an unsaturated surface carbon bond occurs it is unlikely that the adjacent carbon also has an unsaturated bond. Thus in these conditions the growth mode tends to promote the tetrahedral bonding of the diamond structure rather than the planar geometry of the graphite structure. Clearly, in the manufacture of semiconductor devices, it is important to monitor and control surface processes and in particular to achieve an understanding of the kinetic factors that can be exploited to ensure that a complex multilayer device structure maintains its coherence. This control is achieved by the MBE and CVD techniques. We comment here on one small aspect of this field that illustrates the importance of an understanding of surface properties in the design and production of semiconductor devices. A key material of course is Si and at certain stages in device processing it is necessary to `passivate' Si surfaces so that they do not form chemical bonds with molecules in the growth environment. However Si surfaces are very reactive

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because Si adopts the diamond structure in which each atom bonds with four neighbouring atoms (Fig. 1.12) and when a surface is created this leaves the surface atoms with unsatisfied bonds. It is often necessary to passivate such surfaces at various stages in device manufacture. One approach is to deposit an ordered layer of As atoms taking advantage of the natural tendency of As to bond to three neighbours. The difference in the geometry of the Si(100) and (111) surfaces requires different passivating structures. The Si(100) surface is passivated by the formation of As dimers each of which bonds to two Si atoms in the layer below and its neighbouring As in the dimer. On the (111) surface As atoms are coordinated to three Si atoms in the layer below. In each of these structures the As atoms have no unsaturated bonds and the surface is inert ± as illustrated by the fact that the sticking coefficient for the reaction with oxygen is reduced by ~15 orders of magnitude with respect to the corresponding Si surface. This control of surface reactivity is just one of the many ways in which semiconductor device technology is dependent on advances made in surface science.

1.6

Heterogeneous catalysis

It is estimated that industrial processes that employ catalysts have a turnover of the order of £2,000 billion per year in the UK and quite small improvements in the efficiency of catalytic processes can have a major impact on industrial costs. When this is set against a background in which little is known about the actual mechanisms of catalysis one can understand why this field is one of the strongest drivers for research in surface science. The industrial exploitation of catalysis has already had a major impact on the world and one process in particular, the Haber-Bosch method of ammonia synthesis from a mixture of hydrogen and nitrogen gas, has been called the most important invention of the twentieth century because of its role in sustaining the increase in world population through the production of nitrogen based fertiliser (Fig. 1.13).18 This process was discovered empirically and was first used commercially in 1913. The catalytic mechanism is very complex and in spite of a century of research it is still not fully understood. The Haber-Bosch process involves passing N2 and H2 over Fe surfaces. The catalyst is composed of small porous Fe particles of surface area approximately 15 m2/g. A few percent of aluminium oxide (Al2O3), potassium oxide (K2O) and other compounds are added as `promoters' which increase the activity of the iron though their precise role is unclear. The reaction takes place at about 400 ëC in a total pressure of 150 to 300 atmospheres. Ammonia production shows a complex dependence on the partial pressures of N2, H2 and NH3. The reaction can be summarised as follows: N2 and H2 molecules interact with the surface, dissociate into atoms and chemisorb on the surface. The N atoms then react with H atoms to form NH3, which then desorbs from the surface. The main factors

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1.13 Growth in global consumption of N fertiliser and increase in world population (figure by Bryan Christie, reprinted from ref. 18).

that determine the rate of ammonia synthesis are the N2 dissociative sticking probability and the N atom chemisorption energy. As illustrated by the Haber-Bosch process a catalyst must be able to bind molecules, often dissociating them in the process, allow the bound species to react with other molecules and then allow desorption of the desired reaction product. Heterogeneous catalyst systems usually involve small particles that are dispersed over a support material ± often a metal oxide. The small particles have size of the order of nanometres leading to the term `nanoparticles'. Surface effects dominate the structural features of nanoparticles, which have high free energies and are thus very reactive. The fact that catalytic processes usually take place at high pressures means that there is a `pressure gap' in trying to obtain insight into catalytic processes from experiments in UHV conditions. Another limitation on the applicability of surface science techniques to the study of catalysis is that industrial catalysts usually employ highly reactive and poorly characterised nanoparticles whereas surface science seeks to simplify the complexities of chemical reactions by

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experimental work on well characterised single crystal surfaces. These limitations aside, surface science has provided some insight into catalytic mechanisms.19 As an illustration of the strengths and limitations of the current level of understanding achieved in surface science we consider two important examples: the Haber-Bosch process of ammonia synthesis and the oxidation of CO in catalysts for automobile exhausts.

1.6.1 The Haber-Bosch process Surface science studies of single crystal surfaces in UHV have yielded some insight into the mechanisms of the Haber-Bosch process. The results are summarised in ref. 13 and experiments on different surface planes of Fe single crystals show that the efficiency of the process reduces in the sequence: (111), (211), (100), (210), (110). These surfaces have different morphologies, and close examination shows that the (111) surface of BCC Fe is relatively rough exposing the first, second and third layer atoms while the (110) surface is relatively smooth exposing the first layer of atoms and a fraction of the second layer (Fig. 1.2). However surface roughness alone cannot explain the observed catalytic activities which are actually determined by the nature of the active sites. High coordination sites, as found on (111) and (211) surfaces, promote reactions and a good correlation has been found between coordination and efficiency. This suggests why small particles are more efficient than larger ones since small particles have larger numbers of high coordination sites.

1.6.2 The oxidation of CO in automobile exhausts In recent years catalysis has found increasing application in mitigating the environmental impact of automobile exhausts. A typical catalytic converter in a car exhaust is constructed from a high surface area honeycomb structure which is coated with a thin mixture of porous aluminium oxide, ceria and zirconia that is impregnated with nanoparticles of reactive metals such as platinum, ruthenium, palladium and rhodium. Platinum is used to oxidise hydrocarbons and carbon monoxide gas while rhodium serves to reduce nitrogen-oxide species. Early work seeking to understand the detail of these reactions adopted the approach of characterising the surface in UHV, conducting a reaction at high pressure and then examining the surface immediately afterwards. Important information has been gained by this approach, revealing changes in structure and composition of the surface following the catalytic reaction. However, the need to monitor and understand such processes at high pressure is illustrated dramatically by the observation that whereas Ru is the least efficient metal for promoting the oxidation of CO in UHV, it is the most efficient for promoting this reaction at high pressures. Fortunately, significant progress is now being made both theoretically and experimentally in this field and the resolution of the

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apparent paradox of the pressure dependence of the oxidation of CO by Ru, which is described below, is an excellent illustration of the importance of bridging the pressure gap in the study of surfaces.

1.7

Real surfaces: theoretical advances

In spite of the progress made in surface science in the last forty years, our understanding of the behaviour of surfaces in the real world at normal temperatures and pressures is currently quite poor. However, advances are now being made in the study of real surfaces and much of the impetus for this progress has come from theoretical developments. It is now understood why the capacity of Ru to promote the oxidation of CO is so dependent on pressure. The insights into the resolution of this problem are likely to be generally relevant and are worthy of a brief summary. The key issue is to note that the length and timescales at which processes happen and can be studied experimentally in surface science are very different from those that determine the behaviour of surfaces in the real world. This issue has been reviewed recently by Stampf and co-workers20 who draw attention to three important regiems of length and timescale as shown in Fig. 1.14. The bottom left, microscopic, region of Fig. 1.14 is dominated by the motion of electrons and atoms which occur on timescales of 10ÿ15 seconds and 10ÿ12 seconds respectively. An appropriate length scale is 0.1 nm ± the length over which an atom moves between adjacent atomic sites. This regiem is well described by density functional theory which describes the behaviour of

1.14 The spatial and temporal domains relevant to different aspects of the behaviour of surfaces (reprinted from ref. 20 with permission from Elsevier).

24

Surfaces and interfaces for biomaterials

electrons in the potential environment created by atoms and molecules. Many surface science techniques yield useful information on processes that occur on the length and timescales that characterise this regiem. However, while the behaviour of systems in the microscopic regiem is a determinant of what can happen in the mesoscopic regiem, it is not possible to use results obtained in this regiem to infer what will happen at more macroscopic length and timescales. This is because an understanding of the more macroscopic regiems requires the introduction of approaches from statistical mechanics and thermodynamics. Fortunately, theoretical work in association with relevant experimental results is beginning to make useful links between regiems shown in Fig. 1.14. This work resolves the paradox concerning the behaviour of the Ru surface at high pressure since it makes clear that at high pressure the Ru surface converts to the oxide RuO2 and it is the oxide surface that promotes the oxidation of CO to CO2. The key observation is that at higher temperatures and pressures changes in the relative free energies of the metal and its oxide give rise to an intimate interplay between the progress of the chemical reaction and the morphology and chemical composition of the surface and the surrounding gases. We might expect similar considerations to be important in many other areas of surface science where work in UHV does not lead to a simple understanding of the behaviour of surfaces in the real world.

1.8

Real surfaces: experimental approaches

It is clear from the previous section that an understanding of the behaviour of real surfaces under ambient conditions requires information that can only be obtained using experimental probes that can operate outside of UHV. We saw earlier that scanning probe microscopy and optical probes both have this capability. Since, in general, scanning probes reveal information on the microand nanoscale and are rather slow and optical probes yield an average signal over a macroscopic area and can be rather fast these generic techniques complement each other in both the length and timescale over which they can provide information on surface processes. A good example is the study of Hendriksen and Frenken21 who used STM to observe in situ the oxidation of CO on the Pt(110) surface at semi-realistic conditions of sample temperature 425 K and gas flow of pressure 500 mbar. During the reaction the surface was found to change from a metallic CO-covered surface to an oxide surface. In a similar finding to the Ru case, the Pt-oxide surface was found to exhibit the higher catalytic activity. The potential of optical probes to reveal the dynamics of surface processes at higher pressures is shown by the studies of Rotermund and co-workers22 who used RAM to monitor the oxidation of CO at the Pt(110) surface in both low, ~10ÿ4 mbar, and high, approaching atmospheric, partial pressures of CO and O2. This surface has two states, an inert passive state in which the surface is

Fundamental properties of surfaces

25

`poisoned' by the presence of adsorbed CO and an active oxidising surface covered by the dissociative adsorption of O2. Since the surface is not perfect there will always be a particularly active site which at low pressure will promote local oxidation of CO to CO2. This reaction proceeds exothermically and the energy released locally initiates a reaction-front which, in the low partial pressure regimen, spreads across the surface leaving in its wake the poisoned surface created by the exhaustion of locally adsorbed oxygen. The progress of the reaction-front can be observed in real time by RAM and this technique shows that at relatively low pressures, the surface undergoes an oscillating reaction mediated by variations in the partial pressures of O2, CO and CO2. The oscillations are sufficiently robust that two crystal surfaces exposed to the same gaseous environment will oscillate in phase. It is important to note that in the low pressure regimen the thermal characteristics of these reactions are dominated by the large thermal capacity of the experimental mount that is supporting the Pt crystal. However, as the pressure of the reacting gases is increased, a point is reached where energy is released at the reaction site faster than it can be conducted away. The resulting local increase in temperature accelerates the reaction resulting in a thermal runaway that is terminated only by the local exhaustion of the reacting species. In these conditions the smoothly varying oscillations are replaced by sudden local bursts of activity that have no spatial or temporal coordination. This later mode of activity is likely to be more characteristic of the actual behaviour occurring during catalytic reactions at high pressure and demonstrates the importance of experimental work that can bridge the pressure gap.

1.9

Insight into the biological activity of surfaces

This brief survey of progress in surface science suggests a number of ways in which the characteristics of surfaces are likely to have biological significance. We highlight three areas.

1.9.1 The biological activity of nanoparticles As indicated, nanoparticulate matter is characterised by high free energies and a high concentration of active sites. For these reasons they find application in industrial catalytic processes. However, we should expect that these characteristics also endow nanoparticles with significant biological activity and it is well known that ultra-fine air-borne particulates can have harmful effects, as recently demonstrated by the studies of the damage they cause to the mitochondria in cells.23 As the emerging field of nanotechnology develops it would be sensible to be aware of the implications for the health of the population by the release of novel and biologically active particulates into the environment.24

26

Surfaces and interfaces for biomaterials

1.9.2 Insights into naturally occurring biological processes We have drawn attention to the enormous impact on the world of the application of the Haber-Bosch process for fixing nitrogen and the insight that surface science is providing into the detailed mechanisms of this reaction. However, in comparison with natural mechanisms of fixing nitrogen, the Haber-Bosch process is remarkably inefficient requiring, as it does, high pressures and high temperatures. Natural processes clearly extract nitrogen from the atmosphere and bind it in chemical complexes at normal temperature and pressure. Can surface science provide insight into the mechanisms of natural processes in fixing nitrogen? It is interesting to view the recent advance in the understanding of how Mo complexes fix nitrogen25 in a surface science context. To begin with, the mechanism requires a large molecular complex of nanoparticle dimensions. It also depends on an active catalytic site that shows considerable flexibility during the various reactions that constitute the mechanism of nitrogen fixation. This is consistent with what we have learnt from surface science concerning the importance of free energy, active sites and surface flexibility in promoting chemical reactions. One may speculate that the ultimate success of surface science will be to provide us with a sufficient understanding of catalytic processes that we can escape the restrictions imposed by surfaces and develop ways of constructing tailored active and flexible sites in molecules similar to the complexes that operate in natural processes.

1.9.3 Prospects for the study of biomedical interfaces Surface science has addressed metal, oxide and semiconductor surfaces with an approach involving simple systems studied at the fundamental level with powerful probes. This approach will be beneficial for the investigation of complex biosurfaces. One would anticipate that the important determinants of surface processes such as free energies, kinetic factors and the nature of active sites will have an important influence on the behaviour of biomedical interfaces. Clearly the study of such interfaces would benefit considerably from the development of experimental techniques for studying surfaces in ambient conditions. Some indications of the insight into the behaviour of biomedical interfaces can be obtained from studies of the temporal dependence of proteins deposited on human colostrum immobilised on methylated silicon surfaces from human blood serum.26 This study employed a combination of ellipsometry and antibody techniques to provide a convenient and rapid way to indicate the activation of complement sequences on solid surfaces and facilitated a time-resolved determination of replacement sequences and activation pathways.

Fundamental properties of surfaces

27

1.10 Conclusion In this chapter we have considered some of the fundamental properties of surfaces. We have seen that surface science has had a huge impact on the growth of semiconductor devices and has had some success in the understanding of catalytic processes. We have seen that progress is now being made, both theoretically and experimentally, in the study of real surfaces in ambient conditions. We believe that the insights obtained from the study of surfaces have the potential to make a major contribution to the understanding of biological systems.

1.11 References 1. Wyckoff R W G, Crystal Structures, 2nd edn, Volumes 1±6, Interscience, New York, 1963±1971. 2. Woodruff D P and Delchar T A, in Modern Techniques of Surface Science, 2nd edn, Cambridge University Press, 1994. 3. Zangwill A, Physics at Surfaces, Cambridge University Press, 1996. 4. Prutton M, in Introduction to Surface Physics, Clarendon Press, 1994. 5. Venables J A, in Introduction to Surface and Thin Film Processes, Cambridge University Press, 2000. 6. Attard G and Barnes C, in Surfaces, Oxford University Press 1998. 7. Magonov S N and Whangbo M-H, in Surface Analysis with STM and AFM, VCH, Weinheim, 1996. 8. Martin D S and Weightman P, Surf. Sci. 464 23 (2000). 9. Aspnes D E, Harbison J P, Studna A A and Florez L T, J. Vac. Sci. Technol. A6 1327 (1988). 10. Martin D S and Weightman P, Surface Review and Letters, 7 4 389 (2000). 11. Martin D S and Weightman P, Thin Solid Films, 455±456, 752 (2004). 12. McGilp J F, Surf. Rev. Lett., 6 (3±4), 529 (1999). 13. Somorjai G A, in Introduction to Surface Chemistry and Catalysis, Wiley, New York, 1994. 14. Bowker M and Bennet R A, Topics in Catalysis, 14 1 (2001). 15. Schunack M, Petersen L, Kuehnle A, Laegsgaard E, Stensgaard I, Johannsen I and Besenbacher F, Phys. Rev. Lett., 86 456 (2001). 16. de Rome de l'Isle J B L, Crystallographie (Imprimerie de Monsieur, Paris) 1 379 (1783). 17. Radenovic N, von Enckevoly W, Verwer P and Vlieg E, Surf. Sci., 523 307 (2003). 18. Smil V, `Global population and the nitrogen cycle', Scientific American, 227 76, July (1997). 19. Sinfelt J H, Surf. Sci., 500 923 (2002). 20. Stampfl C, Ganduglia-Pirovano M V, Reuter K and Scheffler M, Surf. Sci., 500 368 (2002). 21. Hendriksen B L M and Frenken J M W, Phys. Rev. Lett., 89, 46101 (2002). 22. Dicke J, Erichesen P, Wolff J and Rotermund H H, Surf. Sci., 462 90 (2000). 23. Li N, Sioutas C, Cho A, Schmitz D, Misra C, Sempf J, Wang M, Oberley T, Froines

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Surfaces and interfaces for biomaterials

J and Nel A, Environmental Health Perspectives, 111 455 (2003). 24. Report of Royal Society and Royal Academy of Engineering on Nanotechnology, July 2004. 25. Yandulov D V and Schrock R R, Science, 301 76 (2003). 26. Tengvall P, Askendal A and Lundstrom I, J. Biomedical Materials Research, 35 81 (1997).

2

Control of polymeric biomaterial surfaces

V H A S I R C I and N H A S I R C I , METU, Turkey

2.1

Introduction

Polymers are long, linear or branched chains of molecular weight 10,000 daltons or higher consisting of one or more basic structural or monomeric units. The flexibility of the chain depends on the stiffness of the individual units making up the chain, to the bonds that link them together, the length of the chain and the presence and nature of the branches. These chains are found naturally both in the environment and in the human body. They can also be synthesized from the monomers in the laboratory to yield macromolecules of the desired properties.

2.2

Preparation of polymers

Polymers are prepared by the reaction of monomers. The method of polymerization could be classified as `condensation polymerization' and `addition polymerization'. Condensation polymers are formed by molecules with functional groups that could condense to form the chain mostly by the release of a small molecule such as water, ammonia, etc. Addition polymerization require unsaturated molecules to become saturated while adding to the chain after activation by ionic or free radical initiators, or high energy radiation. The synthetic process is started by the use of an initiator and is helped by exposure to highly energetic radiation (i.e. UV, gamma), elevated temperatures and accelerators. Biological polymers (such as proteins and polysaccharides) are formed exclusively through polycondensation but require enzymes for the initiation and progression of the chain formation reaction.

2.2.1 Homopolymerization Polymers could form by the chain reaction of a single monomer through a condensation or addition reaction. The resultant compound is a very large

30

Surfaces and interfaces for biomaterials

molecule consisting of a single molecule (could be represented by A below) repeated hundreds or thousands of times. -A-A-A-A-A-A-A-A-A-A-A-A-A-A-A-

or

-[A]n-

This kind of a molecule is generally linear unless higher functionality is available on the monomer. Some biological macromolecules such as polysaccharides could yield branched structures only because of the high number of -OH groups. Actually there are no natural homopolymers consisting of amino acids or nucleic acids to yield proteins (i.e. polyalanine) or DNA or RNA (i.e. polyU) from a single molecule. A typical example of this kind of polymer is polyethylene which consists of many -CH2-CH2- groups added on the end. This group is called the repeating unit, since it is the basic element of the macromolecule that repeats itself to represent the final product. The number of these repeating units is known as the degree of polymerization, P, which is an average value due to the random nature of the progression of the reaction. When the carbon atoms in the chain are asymmetric (like -CHR- and unlike -CH2-) the steric position of the monomer side group (R) gains importance. An irregular organization of the R groups lead to an atactic polymer while an alternating position leads to a syndiotactic polymer and a same side orientation is called isotactic. Crystallinity of polymers, a very important property of the solid state, is a direct result of this preferential orientation of the side groups.

2.2.2 Copolymerization and terpolymerization (alternating, block, random) Polymers can also form when two or more monomers are involved in the preparation of the macromolecule. The order in which these monomers come together is very influential on the properties of the resultant polymer. If there are two different monomers such as A and B they could take places in the chain in a random manner forming a random copolymer of A and B: -A-A-B-A-B-B-A-B-B-B-A-A-B-A-A-AThis is the case with most industrial copolymers where the composition of the monomer mixture is not reflected in the resultant polymer. The structural units of the polymer could be ordered in an alternating manner to yield an alternating polymer: -A-B-A-B-A-B-A-B-A-B-A-B-A-B-A-BAnother way the monomer could be arranged on the chain is by way of combination of segments or blocks of single monomer chains: A-A-A-A-A-B-B-B-B-B-B-A-A-A-A-A-A-

Control of polymeric biomaterial surfaces

31

2.2.3 Branching and crosslinking In principle, monomers are either bifunctional or they have unsaturations which enable them to interact with two molecules. In either case a linear molecule is formed. In the presence of compounds with more than two functional groups (for condensation) or more than one unsaturation (for addition) a branched molecule is formed. The branches might form simultaneously with the formation of the chains but also could be induced after the polymer formation is completed (Fig. 2.1). Crosslinking is similar to branching but the general understanding is that a number of chains are bonded together with groups which might (or might not) be of the same chemical origin (Fig. 2.2). As the number of crosslinks increase the solubility of the polymeric structure decreases. At a critical number all the chains of the polymer are bonded covalently together leading to an insoluble polymer, the fully crosslinked structure.

2.1 Branched polymers are linear polymers with short chains of the same composition attached.

2.2 Crosslinking involves covalent and transient linking of polymer chains and significantly modifies the mechanical and solution properties.

2.2.4 Grafting and blending In addition to the above there are graft polymers where one chain serves as a backbone and relatively shorter chains of different chemistry are attached to this backbone as branches (Fig. 2.3). Blending, on the other hand, is a physical mixture where there is no covalent bond between chains of two or more different types brought together either by melting or dissolution followed by solvent removal leading to macromolecules of various types of chains.

32

Surfaces and interfaces for biomaterials

2.3 Grafting of B oligomers on the poly A backbone.

2.3

The solid state and structure

Synthetic polymers are formed in the reaction media and each chain is initiated and terminated at a different time and as a result each chain has a different length leading to a distribution of chain lengths. This is unlike the biological system where the genetic material and proteins are of known fixed molecular weights because their production is highly controlled. In the chemical system there is no such control. As a result, rather than a single molecular weight, one determines an average molecular weight, the value of which is dependent on the method employed. The most widely used average molecular weights are weight average molecular weight (Mw ) and number average molecular weight (Mn ). Their ratio is called the Heterogeneity Index (HI) or polydispersity: w M HI ˆ  Mn and as the HI gets closer to 1 the molecular weight distribution gets narrower. This ratio is especially important in terms of the mechanical properties because as the fraction of very low molecular weight chains increase they act more as plasticizers that reduce the mechanical strength of the material. It is also an important indicator for the mobility of low molecular weight materials to the surface when the polymeric product is introduced into solvents.

2.3.1 Amorphous and crystalline polymers A glassy polymer structure is visualized as densely packed, entangled, random Gaussian coils (Fig. 2.4). The random coil state remains favourable in the glassy state (Stachurski, 2003). Above the glass transition temperature, Tg, where chains gain mobility, the chains are confined to molecular tubes, within which de Gennes reptation takes place. From NMR studies it is confirmed that the chain is not a simple (smooth), constant curvature random coil, but it is further segmented by folds (on itself). The number of folds per chain is a property of the chain (stiffness and chemistry) and of the temperature. The size of the fold is, to a good

Control of polymeric biomaterial surfaces

33

2.4 A glassy polymer structure.

approximation, independent of temperature and chain length. The segments between the folds are between 10 and 20 covalent bonds long. Within the macroscopic volume of the polymer, so-called `free volume' is present as an equilibrium property of the system at Tg. At sufficiently fast cooling rate from this state to a temperature below Tg, the polymer retains a certain amount of the free volume. The actual amount depends on the temperature from which the quench occurs, the cooling rate, and the type of the polymer and its volume. A glassy polymer is formed because the irregular chain architecture prevents crystallization. On cooling to below Tg, a portion of the unoccupied free volume spontaneously diffuses out, allowing the bulk volume to reduce. In the process, the equilibrium end-to-end distance of chains is compressed. Its relaxation time rapidly increases. The prerequisites for the ability to crystallize is regularity, both in terms of composition and stereochemistry of the polymer. Even then it is not enough for the observation of crystallinity. Long periods of cooling might be necessary for allowing the chain substituents to take the required conformation before they can crystallize. They can not, however, completely crystallize; there are always regions (chain ends or side groups) found not to be crystallized when examined by electron microscopy. The resultant structure is what is called the fringed micelle (Fig. 2.5). The rate of cooling is also important because if the cooling rate is high the polymer might not find time to orientate and crystallize and instead solidify in a glassy form. Polymer molecules in the melt are in the form of coils rather than stretched out chains. During the formation of the crystals the coil is still the Ê and the central form. A typical single crystal lamella thickness is around 100 A polymers arrange perpendicular to the plane of the crystal. This has been shown to be true with polyethylene, polypropylene, and low molecular weight paraffins. Since, however, the chains in the lamellae are not equal in length the surface of the single crystal must appear rough.

34

Surfaces and interfaces for biomaterials

2.5 Fringed micelle structure.

2.3.2 Packing, packing density and free volume A glassy polymer is formed because its irregular chain architecture prevents crystallization and it solidifies with a significant volume unoccupied by atoms, which is called the free volume. If this cooling rate is decreased then better packing with a lower free volume is observed. On the other hand when a crystalline polymer is solidified, the organization of the chains and the atoms are such that the unoccupied volume within the solid is minimal. Thus, the free volume in the amorphous polymer is higher than that in a crystalline polymer. The lower free volume indicates a higher amount of interaction between neighboring chain segments and thus to a mechanically stronger structure. In a fringed micelle structure where there is partial crystallinity, certain regions of the polymer are densely packed with low free volume, while the rest is amorphous with high free volume. In such a case the bulk of the polymer as well as its surface has regions of high and low packing density and is expected to have different responses in a biological medium.

2.3.3 Phase states and phase transition Polymers in a solid state could show different properties depending respectively on their chemical structure and presence of crosslinking, the medium temperature and Tg and Tm values of the polymer, the presence or absence of load, and the rate at which it is applied. As a result of this the polymer could function differently than intended, and one needs to know these property changes before converting the polymer into a product. Rubber elasticity When a solid polymeric compound softens upon increase of medium temperature, it still retains the interactions between the chains of a coil and between coils. Upon increase of temperature the chains in a coil and the coils themselves become more and more mobile with respect to each other and few

Control of polymeric biomaterial surfaces

35

contacts are left intra- and intermolecularly. If the temperature is not high enough to completely melt the polymer or if there are covalent linkages (crosslinks) between the chains, it first responds by deforming and then by regaining the original form when subjected to a short deforming force. A change in the position of coils with respect to each other does not take place. This behavior of the material, represented by a strong deformation under a small deforming force and elastic recovery was first observed with natural rubber and is known as rubber elasticity. Relaxation processes When the polymer is brought above its softening point and if the application of the deforming force is sufficiently long, the chains in the structure have time to reorganize with respect to each other into an energetically more favorable state. In this way, the inner tension that has developed in the system is relieved by the process called `relaxation'. If the viscosity of the system is low, the relaxation is too rapid and no recovery can take place. If the viscosity is high the relaxation time is longer and there is a chance for recovery. The latter type of materials flow if the deformation is prolonged, but they are elastic if the deformation is short and rapid. A material that shows both rubber elastic behavior and relaxation is called viscoelastic.

2.4

Polymer-solvent interactions

Polymers present certain chemical and physical properties when dry, but upon coming in contact with liquids all this changes. Certain regions or phases of the polymer might interact with the liquid more than the rest of the structure leading to a different surface composition than in a dry state. The liquid could also lead to swelling or dissolution of the polymer, changing the properties further. It is therefore very important to understand the behavior of the polymer in a variety of solvents.

2.4.1 Polymer gels, hydrogels Crosslinked polymers by definition are composed of a large number of chains strongly (mostly covalently) bonded to each other. Upon immersion of the crosslinked polymer in a solvent the chains forming the structure separate from each other and become solvated but cannot dissolve away due to the restrictions imposed by the links between the chains. This structure is called a gel. If the solvent in which the structure is introduced is water, then the water-swollen structure is called a hydrogel. There are cases where the polymer is introduced to a `less good' solvent, a solvent in which the chemical structures of the solvent and the polymer are not

36

Surfaces and interfaces for biomaterials

similar, and solvation is not as easy as it is in a good solvent. Under those circumstances, an uncrosslinked polymer swells but it is not possible to separate the chains from each other due to the weak interaction with the solvent. In these cases one still obtains gels even though there are no crosslinks. In the latter case, it is possible to achieve solubility by increasing temperature, but a truly covalently crosslinked system cannot dissolve upon increase of temperature.

2.4.2 Sol-gel transition The sol phase is defined as a flowing fluid, whereas the gel phase is non-flowing and maintains its integrity against external forces. Above the critical concentration (critical gel concentration, CGC) of a polymer, the gel phase appears. The CGC is most often inversely related to the molecular weight of the polymer employed. The development of physical junctions is needed for gelation, which must be sufficiently strong to overcome the dissolving forces of the solvent. Thermoreversible gelation of gelatin and polysaccharides such as agarose, amylose, and amylopectin, cellulose derivatives, carrageenans are well known. At high temperatures, they are assumed to have a random coil conformation. On reducing the temperature, they start to form double helices and aggregates. Most natural polymers form a gel phase on lowering temperature. They are said to have Upper Critical Solution Temperature, UCST. However, aqueous solutions of some cellulose derivatives exhibit reverse thermogelation (gelation upon increase of temperature) and are said to have a Lower Critical Solution Temperature, LCST. Cellulose itself is not soluble in water, however, it becomes water soluble upon introduction of hydrophilic moieties (Jeong et al., 2002). When these derivatives have a balance between hydrophilic and hydrophobic moieties, they undergo sol-to-gel transitions upon change of temperature. The temperature at which sol-gel transition takes place depends on the level and location of the substitution. Upon increasing temperature, water becomes a poorer solvent and polymer-polymer interactions become dominant and a gel forms. Table 2.1 shows the LCST values of some polymers.

2.4.3 Influence of bulk and surface properties Creation of polymers with different bulk and surface properties can be achieved in a variety of ways for applications within a wide range of industries. For polymeric materials, the surface energy value is determined mainly by the chemical structure at the surface. It has been suggested that amorphous, comb like polymers possessing a flexible linear backbone onto which are attached side chains with intermolecular interactions exhibit low surface energy values. Surface energy is influenced by parameters such as roughness, the nature of the

Control of polymeric biomaterial surfaces

37

Table 2.1 Polymers with a LCST in water (adapted from Jeong et al., 2002) Polymer

LCST (ëC)

Poly(N-isopropylacrylamide), NIPAM Poly(ethylene glycol), PEG Poly(propylene glycol), PPG Poly(methacrylic acid), PMAA Poly(vinyl alcohol), PVA Poly(vinyl pyrrolidone), PVP Methylcellulose, MC Hydroxypropylcellulose, HPC Poly(N-vinylcaprolactam)

32 120 50 75 125 160 80 55 30

polymer backbone and the pendant chain. For example, grafted perfluorocarbons with varying lengths influence surface properties in a very significant manner (Barbu et al., 2002).

2.4.4 Influence of composition (monomers, polymer type, additives) Polymer-solvent interactions depend on the chemical properties of the solvent, the ingredients of the polymer and the properties of the medium. Every system tends to reduce its internal energy (
U) or enthalpy (
H) and increase its entropy (
S). In other words every system tends to decrease its Gibbs Free Energy (
G). The relation between free energy, enthalpy and entropy at a certain temperature is given as:
G ˆ
H ÿ T
S When
G is decreased a process is spontaneous. During the dissolution process there is an increased mobility of the solute molecules and thus there is an increase in entropy. Enthalpy is also expected to decrease because there should be increased interaction between the solute and the solvent molecules if there is a good solvent for the solute. On the whole, free energy is decreased (
G < 0), making the dissolution process spontaneous. The enthalpy change is an outcome of the solvent-solute interaction. When there are more and stronger interactions leading to decreased enthalpy, this leads to increased solubility. The major factor at this stage is the chemistry of the monomer chains and of the resultant polymer. For example, if the polymer is composed of hydrophilic monomers such as N-vinylpyrrolidone, than the interaction with water will be strong and the enthalpy will decrease, leading to decreased free energy and increased solubility. If, however, the molecule in question is styrene, then the polymerwater interaction will be less than polymer-polymer and water-water interactions. Thus, the result would be insolubility. Only upon increase of

38

Surfaces and interfaces for biomaterials

temperature is mobility of chains increased leading to higher entropy and thus eventually a temperature is reached where solubility is achieved.

2.5

The polymeric surface and surface-bulk difference

Polymers are synthesized to yield glassy solids or powders. It is not easy to distinguish between the surface and the bulk. Upon conversion of the polymer into a product, the bulk and surface become different both in terms of chemistry and topography. There are a variety of reasons for the bulk and surface chemistry to differ: · oxidation of the surface · orientation of the macromolecules in a way that gives better interaction with a mould · different free energies of the various components of the macromolecule. Upon introduction into a liquid medium, the differences become more distinct. The interaction of the polymer chains with the solvent modifies the surface and bulk in terms of concentration and localization of hydophilic and hydrophobic groups. Upon introduction to a hydrophilic liquid, the concentration of hydrophilic groups or segments on the surface increases while in the interior hydrophobic groups concentrate. As a result, the surface properties of the polymer become different from those of the bulk. The packing densities on the surface and in the bulk are different. Due to thermodynamic constraints, the hydrophobic groups and segments form more hydrophobic interactions that lead to an increase in the packing density while on the surface, interaction between the solvent and polymer overcomes the polymer-polymer attraction leading to a swollen or highly extended conformation, and a lower packing density. In good solvents, the polymer-solvent interaction is more favorable than the polymerpolymer one. Polymer chains in good solvents swell due to steric repulsion. This spatial size of a polymer coil is much smaller than its extended contour length but larger than the size of a typical chain. The reason for this peculiar behavior is entropy combined with the favorable interaction between polymers and solvent molecules in good solvents. Similarly, for the adsorption of polymer chains on solid substrates, the conformational degrees of freedom of polymer coils lead to salient differences between the adsorption of polymers and small molecules. In the case of `poor' solvent conditions, the effective interaction between polymers is attractive, leading to collapse of the chains and to their precipitation from solution (phase separation between the polymer and the solvent). In this case, the polymer size decreases, like any space filling object embedded in threedimensional space. Thus if the polymer is a homopolymer with no segments of different functionalities, then in the poor solvent the packing on the surface is expected to increase in comparison to the chains in the inside.

Control of polymeric biomaterial surfaces

2.6

39

The general properties of a biomaterial surface

In the past, biomaterial biocompatibility was considered to be passive behavior towards the biological system. This requires the biomaterial to be nonthrombogenic, non-allergenic, non-carcinogenic, and non-toxic (Klee and HoÈcker, 2000). Williams defined biocompatibility, however, as `the ability of a material to perform with an appropriate host response in a specific application' (Williams, 1999). A variety of properties determine biocompatibility. Among these are the mechanical and chemical/physical properties of a material, both in the bulk and at the surface; the surface exerts the heavier influence because it is the component where the biomaterial and the biological system meet and interact. The properties of the surface that have utmost importance are: · the chemical structure ± hydrophilicity ± presence of groups that could initiate reactions in the biological system · the morphology ± the distribution and abundance of hydrophilic/hydrophobic and crystalline/amorphous phases ± surface topography, i.e. surface roughness and the presence of physical forms. The surface characteristics of a polymer are considerably different from the bulk characteristics. In a dry state, due to the minimization of surface energy and chain mobility, the non-polar groups move to the phase boundary with air while the reverse happens in an aqueous medium. Following this, low molecular weight components could migrate either towards or away from the surface leading to differences in the properties of the surface and the bulk. At the phase boundary between the biomaterial and the aqueous surrounding tissue, a different situation arises than at the phase boundary between the biomaterial and air. When the implant comes into contact with the biological system the following reactions are observed: 1.

2.

Within the first few seconds molecules (especially proteins and lipids) from the surrounding body liquids and tissue are deposited. In blood the properties of this adsorbed protein layer determines hemocompatibility, while in tissues it controls further reactions of the cell system. The nature of the adsorbed proteins is dependent on the surface characteristics of the implanted material. Upon implantation, the tissue neighboring the implant undergoes mechanical, chemical or physical damage or a combination of these. This damage could be acute (short-term and intense) or chronic (long-term) depending on the degree of stability of the material surface. In any case a period of inflammation is observed.

40 3. 4. 5.

Surfaces and interfaces for biomaterials A biocompatible implant, if inert, is surrounded by a thin layer, mainly consisting of collagen. A biocompatible implant, if not stable will undergo some sort of degradation and/or erosion in the harsh biological medium. This will continue until the implant is completely removed from the medium. A nonbiocompatible implant, if not stable and inert, will undergo degradation and erosion but its products will initiate a variety of undesirable reactions.

Since the surface is the outermost region of the implant interfacing tissue, it should fulfill the requirements of biocompatibility.

2.7

Modification of polymer surfaces

Surfaces of polymers are modified intentionally and unintentionally by various mechanisms and approaches as described below.

2.7.1 Oxidation by air Oxidative degradation is a route for unintentional surface modification and plays an important role in the aging of polymers. Since mainly the surface of a polymer is exposed to air, the effect of oxygen is observed more on the surface than the bulk. This oxidation is influenced by light, water, moisture and temperature. With most polymers, the exterior signs of oxidation are yellowing and if oxygen penetration was possible, then as an increase in brittleness. Unsaturated polyolefins are susceptible to oxidation by air. Oxidation generally induces the formation of peroxide which later leads to chain scission and probably simultaneous free radical formation. This decrease in chain length causes a deterioration of mechanical properties. Thus upon exposure to air chain length decreases and surface oxidation leads to more polar groups than initially. The extent of this change in the bulk is very limited and is controlled by the gas permeability of the solid polymer.

2.7.2 Grafting Grafting can be achieved in a variety of ways leading to products with different forms and chemistries, such as interpenetrating, crosslinked, brush, stimuli responsive, etc. Interpenetrating graft Interpenetrating graft copolymers are obtained where a polymer is dissolved in the monomer to be grafted and then allowing polymerization to take place. The

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chains of the earlier polymer will be entrapped in the newly formed structure. If this new polymer has functional groups or more unsaturations than one, the newly formed structure will be crosslinked within itself more strictly entrapping the original polymer. When surface modification via crosslinked grafts is considered, then the approach would involve wetting or swelling of the surface of the polymer by the monomer and then allowing polymerization to take place. The thickness of the grafted region will vary with the extent of wetting of the surface before initiation of polymerization. In this manner the surface of a mechanically suitable hydrophobic polymer could be rendered hydrophilic for increased hemocompatibility or for decreased lipid adsorption. Crosslinked graft Crosslinked graft copolymers are obtained where the unsaturated polymer is dissolved in the monomer to be grafted and allowing polymerization to take place. The newly formed chains could terminate by a reaction with the unsaturations on the chains and lead to crosslinkages. When surface modification via crosslinked grafts is considered then the approach would be as in the case of the interpenetrating graft. Brush graft Brush grafting could be achieved by activating the functional groups on the polymer to react with newly introduced monomers leading to brushes on the surface. For example, in a typical application, hydrophilic polymer brushes bearing alcoholic hydroxyl groups were introduced onto a polyethylene (PE) surface by radiation-induced grafting of 2-hydroxyethyl methacrylate (HEMA), vinyl acetate (VAc), and glycidyl methacrylate (GMA). HEMA already carried a hydroxyl group. VAc or GMA grafted membranes, on the other hand, were hydrolyzed to yield the hydroxyl groups. Thus, a polyethylene surface was made hydrophilic by brush grafting. The brush presence also modified the properties of the material. When adsorption of gamma globulin onto this material was tested it was observed that upon increase of the amounts of the hydrophilic polymer brushes, adsorption of gamma globulin on the PE base was decreased. Also, these hydrophilic polymer brushes bearing multiple glycol groups (from GMA) and alcohol groups (from VAc) provided an attractive site for the linkage of ligands using chemical methods, a property which was not available for untreated PE (Kawai et al., 2003). Stimuli-responsive grafts Stimuli-responsive polymers respond to small changes in their environment with responses that can be used in various applications (Jeong and Gutowska, 2002).

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The most frequently applied changes or stimuli are temperature, pH, and ionic strength. The responses of polymers grafted on surfaces are presented when introduced into solutions. Recent advances in the design of stimuli-responsive polymers have created opportunities for novel biomedical applications. The responses can be observed as shape, surface characteristics, solubility and as enabled sol-gel transition. The extent of the response needs to be controlled, and therefore the reasons (mostly thermodynamic) for them have to be understood. A typical, and most commonly used, stimulus (temperature) responsive polymer is NIPAM (N-isopropyl acrylamide). NIPAM is soluble below 32 ëC and precipitates above 32 ëC in water. Below the phase transition temperature, i.e., LCST, the hydrogen bonds between the polymer and the water molecules are favored, and thus the polymer stays soluble. The chains are fully extended. Above this temperature the hydrogen bonds are ruptured due to excessive thermal motion of the molecules and the polymer chains prefer to interact with each other leading to insolubilization. The properties of this polymer could be modified by copolymerization with other monomers. Copolymerization of NIPAM with butylmethacrylate for example, decreases the LCST of aqueous copolymer solution because the co-monomer is hydrophobic. Copolymerization with hydrophilic co-monomers on the other hand results in an increase in LCST thus enabling modified responses of a system constructed of NIPAM. If two different stimuli-responsive hydrogels are brought together (i.e. NIPAM and Nvinyl caprolactam) in a hydrogel then two different response ranges (such as two different inversion temperatures) can be obtained. The changes in chain extension levels are also important in certain tissue engineering applications. When the temperature of the medium was decreased below 32 ëC, confluent cardiac myocyte layers were lifted from cell culture dishes without need for trypsinization, an enzymatic treatment which might be harmful to the resultant cell culture if not used properly. The sol-to-gel transition temperature of PEG±PLGA±PEG triblock copolymers was shown to be controlled over a temperature range of 15 ëC to 45 ëC in aqueous solution by changing PLGA and PEG length and the ratio of lactic to glycolic acid (Jeong and Gutowska, 2002). Owing to the triblock topology, PEG±PLGA±PEG polymers have limitations in terms of molecular weight and degradation profile and the ability to show the sol-to-gel transition in a desired range of ~10 ëC to 30 ëC. Graft copolymers of PEG±g±PLGA and PLGA±g±PEG, however, present sol-to-gel transitions at ~30 ëC. PEG±g±PLGA copolymers have hydrophilic backbones and form gels with short durability, whereas PLGA±g±PEG copolymers have hydrophobic backbones and form much more durable gels. Poly(L-lysine)±g±poly(histidine) is an example of a pH sensitive or responsive system (Jeong and Gutowska, 2002). Poly(L-lysine) is positively charged at physiological pH. The pKa of poly(histidine) is 6.0, and it therefore undergoes conformational changes at pH lower than this. Poly(propyl acrylic

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acid) (PPA) and poly(ethacrylic acid) (PEA) are also sensitive to pH changes in the environment. Both polymers have the capability to ionize at high pH and are neutral at lower pH. Thus, they are soluble at high pH and insoluble or less soluble at lower pH. This is reflected in the form of abrupt conformational changes at pH around 5 to 6. Certain polymers have the capability to hypercoil or to form hydrophobic bonds to create compact molecules, and this makes it possible to induce the macromolecules to change their conformation in response to local stimuli. Polymers with weakly charged pendant groups, i.e., either weak acids or bases, exist in the form of extended chains due to repulsion between these charged groups. If, in addition, the polymer also bears alkyl or aromatic pendant groups, then the latter will initiate hydrophobic interactions which will lead to the hydrophobic groups being located in the interior of the structure, thus enabling maximal hydrogen bonding between the polar groups and the water molecules. This `hydrophobic effect', is also the principal driving force in the formation of lipid-based assemblies such as cell mebranes and in determining the conformation of native proteins. Once medium properties are altered, the conformation of the molecule can also be reversibly changed (Tonge and Tighe, 2001).

2.7.3 High-energy treatments Polymer surfaces can be activated by application of high-energy radiation such as glow discharge plasma or gamma irradiation. Plasma modification Low-temperature plasmas are produced by electrical discharge in low-pressure gases. They consist of a mixture of highly reactive species, i.e., ions, radicals, electrons, photons and excited molecules. Their chemical composition and physical characteristics are determined, in addition to the gas used, by device parameters, such as chamber geometry, gas flow rate, frequency and the power applied. This method is used to modify the chemistry and morphology of polymer surfaces to a depth of several tens of microns thus leaving the bulk properties practically intact. Thus, completely different chemistry, hydrophilicity or surface roughness can be obtained. Riccardia et al. (2003) used air as the gas to achieve plasma treatment of polyethylene terephthalate. The surface chemical and physical modification on PET fibers was a remarkable increase in hydrophilicity, extensive etching and a low molecular weight. XPS studies revealed formation of C±O and C±N bonds in the surface layer and the simultaneous decrease of C±C and C±H bonds. They interpreted etching as mainly a consequence of ion bombardment, while surface chemical modifications were mainly due to the action of neutral species on the

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plasma-activated polymer surface. The main effects of the active species created from the gas on polymer chains is mainly chain scission with new functional group creation and crosslinking. Grafting of nitrogen-carrying functional groups by plasmas requires nitrogencontaining gases. This is achieved through generation of nitrogen-containing radicals or excited N2, atomic nitrogen, NH, NH2 radicals, etc. Grafting of amino groups by plasmas is generally known to occur in combination with that of other nitrogen functionalities (Meyer-Plath et al., 2003). In another study polyethylene (PE) surfaces were treated with corona discharge with power ranging from 10 to 50 W (treatment duration 5 s) using air as the gas; the wettability of the PE sheet was significantly increased as judged by the water contact angle. New C±O, C=O and O±C=O bonds were detected with XPS (Lee et al., 2003). Gamma irradiation One of the most common irradiation types in industrial use is gamma. This is mainly produced by a Cobalt-60 source. Their main advantage is that they are very penetrating. Crosslinking and chain scission are the two major effects that gamma radiation has on polymers. This exposure could be applied for the purpose of processing or for sterilization. Radiation grafting and hydrogel formation are among the more important gamma irradiation applications (Clough, 2001). Grafting is employed because with it the surface properties can be tailored according to needs, while the material retains its bulk properties. Some outcomes are an improvement of chemical resistance, wettability, biocompatibility and hemocompatibility, dyeability of fabrics and antistatic properties. They could also be made to attach functional groups to enable the material to immobilize enzymes and other bioactive species. The advantage of radiation crosslinking over chemical crosslinking is that the former is more economical, faster, allows substantial decrease in the chemicals in the material processed (thus lower risk of bioincompatibility or allergenic response), and can be carried out at around room temperature. When irradiated, polymers are either inclined to undergo crosslinking or chain scission depending on the chemistry and the processing conditions. A polymer that undergoes chain scission could crosslink when the conditions are modified. Alterations in surface morphology can occur upon irradiation, especially in the form of an increase in roughness, a change that alters properties such as contact angle and cell adhesion.

2.7.4 Self assembled monolayers (SAMs) Self assembled monolayers are spontaneously formed highly ordered structures guided by the surface on which attachment takes place, by the components of the macromolecule and of the solvent. A well known example is alkane thiols on

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gold surfaces where the thiols attach to the gold surface with alkane tails sticking out and eventually align to form a tightly packed crystalline structure. The properties of this assembled layer depend very much on the alkane length. PEI (polyethyleneimine)-PLGA block copolymers were also found to be selfassembled in water where the PLGA segment acted as a hydrophobic aggregate block and the PEI segment as a hydrophilic corona-forming block (Nam et al., 2003). The block copolymers formed micelle-like aggregates in water and the size of the aggregates depended on hydrophobic block length and the ionic state of the hydrophilic block. The aggregate size decreased when the PLGA block length decreased and the PEI block was protonated. The amphiphilic nature of block copolymers consisting of hydrophilic and hydrophobic blocks provides an opportunity to form micelle-like structures in water. However, they do not form if they are introduced directly to an aqueous medium. They form their micellelike aggregates only when the solution of the polymer in an organic solvent such as DMF is introduced into aqueous media and then subsequently dialyzed, allowing time for organization in the form of micelles.

2.7.5 Surface patterning The use of chemical and physical patterns on polymer surfaces is on the rise due to the variety of uses they can be put to. Among them are biomedical uses such as construction of biosensors and tissue engineering applications. Patterning can be achieved by various methods as described in the following paragraphs. Contact printing/microprocessing This is basically a photolithographic approach. A variety of processes such as embossing, fused deposition modelling, ink jet printing, microcontact printing, proton micromachining and rapid prototyping are among the methods that could be classified under this heading. Embossing The first step is to fabricate a master die using electron beam lithography from which all other devices will be made. This master is made using a quartz substrate which has been coated with a thin metal film layer of Ti/Pd/Au to remove charge during exposure and to act as a plating base (Casey et al., 1997). A negative e-beam resist which facilitates the writing of small features over a large area in a relatively short time is used. The resist is developed to produce the pattern which is intended to be transferred into the plastic. The sample is then cleaned using an organic solvent to remove any remnants of resist or primer. It is now electroplated with nickel to the required depth which is typically a half to a third the height of the resist. The resist is removed by

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refluxing in a solvent after which the sample is a nickel die ready to use as an embossing master. If the features required for embossing are less than 150 nm in size, then the master is made by dry etching the resist pattern into quartz using an intermediate titanium layer. The quartz is then etched. The polymer used for embossing could be cellulose acetate which is in l mm thick sheet form. The important characteristics of this plastic are that it has a smooth, gloss surface with good clarity and a relatively low flow temperature in the range 135±175 ëC. The embossing process requires application of a constant force to the die/plastic/ platen stack. It is essential that a constant pressure is maintained throughout not only at the imprinting stage of the procedure, but also during the cooling of samples. Standard embossing times are 30 minutes at the elevated temperature of 135 ëC, followed by a 15-minute cooling period. Fused deposition modelling (FDM) process The FDM method forms three-dimensional objects from computer-generated solid or surface models like a typical rapid prototyping process. Models can also be derived from computer tomography scans, magnetic resonance imaging scans or model data created from 3D object digitizing systems. FDM uses a small temperature-controlled extruder to force out a thermoplastic filament material and deposit the semi-molten polymer onto a platform in a layer by layer process. The monofilament is moved by two rollers and acts as a piston to drive the semimolten polymer. At the end of each finished layer the base platform is lowered and the next layer deposited. The designed object is fabricated as a threedimensional part based on the deposition of thin layers of the polymer. The deposition path and parameters for every layer are designated depending on the material used, the fabrication conditions, the applications of the designed part and the preferences of the designer. This method was employed in the production of novel scaffolds with a honeycomb-like pattern, fully interconnected channel network, and controllable porosity and channel size (Zeina et al., 2002). Poly(-caprolactone) (PCL) was developed as a filamentmodeling material to produce porous scaffolds, made of layers of directionally aligned microfilaments. The PCL scaffolds were produced with a range of channel sizes 160±700 m, filament diameter 260±370 m and porosity 48± 77%, and regular geometrical honeycomb pores, depending on the processing parameters. Ink jet printing technique The technique is very simple. The polymer is dissolved in a volatile solvent (e.g. chloroform, trichloroethylene) and the liquid is ejected by the printing head on the substrate (i.e. alumina). As the drop reaches the surface the solvent evaporates leaving the solidified polymer behind. If more than one drop is

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ejected the solidification time is increased, leading to wider, thicker and more homogeneous polymer deposition (Pede et al., 1996). Microcontact printing of DNA In this method a rubber stamp that carries the inverse of the required pattern is immersed in the patterning solution (DNA, protein, enzyme, antibody, etc.) and then applied on the surface of the target material. The pattern is generally produced by photolithography on silicone based rubbers. In a typical application, photoresist patterns of 5 m line and 10 m space were prepared by photolithography on a silicon substrate which was used as a master for the stamp (Fujita et al., 2002). Patterned poly(dimethylsiloxane) (PDMS) stamps were made by mixing silicone elastomer and a curing agent and pouring the mixture in a plastic disposable container with the silicon master at the bottom. This was then heated in an oven, and the hardened polymer removed. A dilute solution of 3-(2-aminoethyl amino)propyl triethoxy silane (amino silane) was applied directly to the stamp and was microcontact printed on mica substrates. DNA solution was applied on the patterned substrate, washed with pure water, and dried under nitrogen gas thus creating a pattern of DNA on the mica surface. Proton micromachining In this application a beam of protons is focused on the target and scanned to create the pattern needed. A resolution of 1 m could be achieved by this approach. Sanchez et al. (1999) used a piece of PMMA as the template material, exposed to a 0.6 MeV proton beam using a nuclear microscope. The beam spot was scanned over the PMMA in a series of patterns consisting of nine different types of ridge/groove structures. Multiple repeated exposures of the patterns were carried out to achieve a homogeneous exposure. Rapid prototyping This system involves fabrication of scaffolds for tissue engineering applications via robotic desktop rapid prototyping. In one recent application the experimental setup consisted of a computer-guided desktop robot and a pneumatic dispenser (Ang et al., 2002). As dispensing material, chitosan and chitosan/hydroxyapatite (HA) dissolved in acetic acid was chosen. It was forced out through a small nozzle into a dispensing medium which was sodium hydroxide and ethanol. Layer-by-layer, the chitosan was fabricated with a preprogrammed pattern. Neutralization of the chitosan formed a gel-like precipitate and the hydrostatic pressure in the sodium hydroxide (NaOH) solution kept the cuboid scaffold in shape. A good attachment between layers allowed the chitosan matrix to form a fully interconnected channel architecture.

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Laser etching (laser photoablation) Laser photoablation or laser etching is a highly efficient and versatile etching process by which selected regions in a polymer layer can be removed by highenergy UV laser irradiation. The material is removed in a one-step process, requiring no developing or other steps before or after exposure. In addition to a wide variety of polymeric materials, photoablation has also been successfully used for laser-etching of several metals, oxides and other inorganic materials. The largest applications of photoablation include drilling of vias in multilayer microelectronic circuits, fabrication of nozzles in inkjet printheads, and corneal shaping for vision surgery. Numerous other applications of ablation have been implemented and many others are being researched in a variety of electronic, optoelectronic and medical fields. In the laser ablation (LAB) process using high-energy ultraviolet laser beams, polymers are chemically decomposed and removed without heating, therefore, smooth bottoms and sidewalls of the etched holes can be obtained. In practice, a polymer film is glued on a Si substrate and is etched by the LAB process. Slant planes can be easily formed by the relative movement of the laser beam. In a typical system the light source is a KrF excimer laser operating at a wavelength of 248 nm. The laser beam radiating from the source goes through a mask and a pattern on the mask is imaged on a sample. The sample is etched vertically in the irradiated area and the depth of the etched hole is proportional to the number of laser pulses. Laser photoablation can be carried out either in a spot-by-spot fashion using a focused beam from a low-power laser and raster-scanning it to address all the locations, or by projection-imaging the desired pattern of vias, lines, etc., from a mask onto the substrate using a high-power laser, thereby ablating thousands of features simultaneously. In the case of near-IR (NIR) laser patterning, however, it is believed that laser etching mostly results in substantially disordered profiles due to thermal effects. It is accepted that the thermo-mechanical properties of polymer materials play an important role in determining the quality of NIR laser etched surfaces. E-beam etching Conventional lithography involves patterning with light, generally in the UV region, of a thin polymer film that is spin-coated onto a silicon wafer. The radiation crosslinks (negative resist) or degrades (positive resist) the exposed polymer which is then washed (developed) to remove or retain the design, leaving behind a patterned polymeric surface. The same approach could be applied to a polymer with a significant thickness to create a pattern on the surface while retaining the bulk properties. This could be achieved by e-beam irradiation. The e-beam machine plays a significant role in the processing of polymers in a similar fashion. A number of different designs and energies are available.

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Industrial e-beam accelerators with energies in the 150±300 keV range are used when low penetration is needed. Accelerators operating in the range 1.5 MeV are used where more penetration is needed. High-energy commercial e-beams (10 MeV) are used in the sterilization of medical supplies already boxed and ready for shipment. They have high dose rates and therefore short processing times. Low penetration indicates more effective use of e-beam energy than with gamma. In patterning 100 keV electron beams a magnetic lens is used. Patterning is achieved by using an electron scattering mask. E-beam is capable of patterning at around 70 m resolution (Clough, 2001).

2.7.6 Immobilization of functional groups and molecules Chemical modification of the surface through introduction of new groups to the polymeric product is designed to impart certain properties to a material. This does not necessitate the modification of the bulk. In a typical application, thiocyanation of plasticized PVC led to a product with increased hydrophilicity but also significantly reduced retention of two types of bacteria strongly implicated in implant related infections, S. aureus and S. epidemidis (Nirmala and Jayakrishnan, 2003). Immobilization of heparin onto the surface of biomaterials to render them hemocompatible has been used in a large number of studies. Heparin immobilization was reported with PVA, methyl methacrylate, and polyurethane to name a few. Heparin was immobilized on Gore-TexÕ to improve its vascular graft performance (Begovac et al., 2003), to introduce hemocompatibility to nonwoven fabrics of polypropylene (Tyan et al., 2002), to silicone rubber, polyethylene, polypropylene and polyvinylchloride (Michanetzis et al., 2003), to PLA (Zhu et al., 2002), etc. Another approach is surface molecular imprinting. This technique helps introduce specific recognition sites into polymers. It has the ability to be applied to a wide variety of target molecules. If a high interfacial activity functional molecule is used, the enzyme-mimic site can be located on the surface of the polymer. In a specific example, the technique was used to help resolve optically active amino acids (Toorisaka et al., 2003). 2-Methacryloyloxyethyl phosphorylcholine (MPC) polymers were synthesized to mimic the biological membrane structure. MPC polymers are useful for surface modification of conventional materials even when random copolymers composed of MPC and alkylmethacrylate are applied as coating polymers. They effectively reduce protein adsorption and denaturation and inhibit cell adhesion even when the polymer is in contact with whole blood in the absence of any anticoagulants. It was observed that phospholipid polymer surfaces showed excellent blood compatibility (Yamasaki et al., 2003).

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2.8

Surface analysis

A biomaterial surface presents different levels of crystallinity, roughness and chemical groups to the biological environment into which it is introduced. Immediately after implantation, the adsorption of a variety of biological molecules such as proteins and lipids change the surface further (Fig. 2.6). In order to be able to construct the ideal biomaterial, the surface has to be characterized and this is achieved by a variety of techniques. In order to determine the composition and structure of a biomaterial surface, different methods which provide varying degrees of information are commonly used. ATR-IR (attenuated total reflectance infra-red) or ATR-FTIR (attenuated total reflectance Fourier Transform infra-red) spectroscopy supplies the characteristic absorption bands of functional groups with an informational depth of 0.1±10 m. Samples with rough surfaces are studied with photoacoustic spectroscopy (PAS), which allows analysis down to approximately 20 m. The achieved informational depths are usually larger than the thickness of the modified interface which might have a several molecules thick modification as in-plasma modification approaches and therefore the spectra obtained could include peaks due to the bulk as well. X-ray photoelectron spectroscopy (XPS) (or ESCA, electron spectroscopy for chemical analysis) is a more surfacesensitive analytical method which supplies information not only about the type and amount of elements present but also about their oxidation state and chemical surroundings. Depths of approximately 10 nm can be achieved with this method (about 50 atomic layers). A more surface-sensitive method is secondary ion mass spectroscopy (SIMS) where primary ions interact with the polymer surface and the mass spectra of the formed ions (secondary ions) are obtained which give information about the chemical composition of the outermost atomic layers (approximately 1 nm in thickness). The application of atomic force microscopy (AFM) in comparison to scanning electron microscopy (SEM) delivers information about surface properties as far as molecular dimensions. Another advantage of AFM compared with SEM is that the sample is investigated in the original state (no sputtering).

2.6 A typical biomaterial surface in the in-vivo environment.

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The full characterization of the surface of a biomaterial is often verified only by application of several analytical methods to the sample. The depth of the various analytical systems could deliver information about are: SIMS and AFM about 1 nm, XPS about 10 nm and ATR-IR 4000 nm.

2.8.1 Morphology When modified in one of many ways, the morphology of a polymeric surface is also changed whether intended or not. In order to be able to study the extent and type of these changes, a variety of novel techniques are used. The following three microscopic techniques are among the most effectively and extensively used surface characterization techniques. Scanning tunnelling microscope (STM) Two researchers working at IBM in Switzerland invented in 1982 the scanning tunnelling microscope (STM), for which they were awarded the Nobel Prize for Physics. In this technique, a fine sharpened tip, which is presumed to have a single atom at the apex, is scanned above the surface of the sample. A voltage (bias potential) is applied between the tip and the sample and a small current (tunnelling current) which flows between the gap is measured. The variation in tunnelling current as it passes over the atomically corrugated surface is then recorded and, if this process is repeated across the entire sample, a threedimensional map of the surface can be obtained. The choice of materials investigated was partly limited by the requirement for them to be conducting, so that a tunnelling current could be measured. Atomic force microscopy (AFM) Atomic force microscopy (AFM) is a method increasingly being used to study and quantify surface properties of materials in their untreated, natural form. It is essentially a scanner that creates topographical maps of surfaces. A very sharp tip, located at the free end of a cantilever follows the contours of the surface as it is moved over it. The deflections of the tip are measured by an optical detector and recorded by a computer. This instrument is very similar to STM, except that the tip and the surface are in contact and interatomic van der Waals forces acting between them provide the contrast mechanism, rather than a tunnelling current. There was no longer a requirement for the sample to be conducting, thus opening the way for the study of many other materials (Smith et al., 1997). The data can be used to create a 2-D or 3-D image and measure surface roughness. It is very important in the biomedical sciences because most polymers are highly hydrated in vivo, and thus treatments for visualization as in SEM (high vacuum, conductive coats, etc.) would not maintain their actual surface

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topography. With AFM, however, it was possible to examine very fragile hydrogel materials like soft contact lenses in their swollen form and measure their surface roughness (Grobe et al., 1996). It is also possible to monitor real-time events like hydrolytic degradation. In a study by Davies et al. (1996) the degradation of the components of blends of poly(sebacic acid anhydride) (PSA) and PLA was studied revealing the preferential degradation of the PSA component leaving behind surfaces enriched in PLA. It was thus possible to expose PLA morphology. With the use of a similar technique, scanning force microscopy, close packed, needlelike organization of crystals of poly(butene-1) were shown with a resolution of the order of nanometers (Jandt et al., 1993). It was claimed that individual poly(butene-1) molecules could be observed. In another study, lamellar organization of polyethylene (Jandt et al., 1994) and various spherulitic surfaces (Lustiger et al., 1989; Harron et al., 1996) could be visualized. A similar observation was made by Hasirci et al. (2002) where degradation of PPF-NVP/ EGDMA reinforced PLGA was investigated revealing a lamellar organization of the bone reinforcement plates after a very brief incubation in distilled water. Scanning electron microscopy (SEM) SEM is a microscopic method extremely valuable in polymeric products and biomaterials areas along with many other research and application areas. Its main attributes are that it reveals the 3D topography of the specimen examined and its magnification could be extremely high (factors of hundred thousands). For polymeric materials, high magnifications are generally not possible because of the intensity of the electron beams damaging the thin polymeric samples which generally lead to deformation and even melting of the specimen. Other limitations of SEM are the need for a high vacuum which prevents visualization of solvated samples in their natural state and the need to coat the sample with conductive materials such as gold. In order to overcome these problems, methods of fixation or stabilization and obtaining durable replicas of the actual sample are tried with varying degrees of success. The resultant 3D image is generally worth the effort because it reveals a wealth of information about the specimen examined. SEM was used in a series of studies on drug delivery systems and bone plates with great success. A PLGA rod (or rather fiber with diameter ca. 0.65 mm) controlled pain relief system was designed to deliver analgesics for a duration of 3±4 weeks (Hasirci et al., 2003). The initial tests were carried out in rats by tying the fiber to the sciatic nerve before testing for the loss of sensitivity to external pain inflicting stimuli such as high intensity light. The fiber withstood implantation very well, without disintegrating with the only deformation being at the location where it was tied to the sciatic nerve. The PLLA-PPF based bone plates for fracture fixation in rats reveal that while one would perform its function of stabilizing a fracture properly the other implant would probably disintegrate by cracking under the applied stress (Hasirci et al., 2000).

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2.8.2 Chemistry The surface chemistry of a polymer intended for biomedical applications can be investigated by a number of approaches, mostly based on spectroscopy and surface energetics. In the following section, information about these methods is provided. X-ray photoelectron spectroscopy (XPS) ESCA is the abbreviation of electron spectroscopy for chemical analysis, and is the same as XPS (X-ray photoelectron spectroscopy). It is a method for studying the energy distribution of electrons ejected from a material that has been irradiated with a source of ionizing radiation such as X-rays. This powerful tool provides quantitative information about basic properties such as binding energy, charge, and valence state, which examines the atom as a part of its chemical environment. The impact of ESCA in polymeric characterization has been twofold: it can analyze relatively intractable materials without the need for special sample preparation and it is a surface sensitive method. In principle, when any material is bombarded by photons with energy greater than the binding energy of an electron in a given atomic shell or sub-shell, there is a finite probability that the incident photon will be absorbed by the atom and an electron is either prompted to move to an unoccupied level or ejected as a photoelectron. This depends on the energy of the incident photon and the atomic number of the target element. The kinetic energy of the photoelectron is: KE ˆ h ÿ BE where KE and BE are the kinetic energy and the binding energy of the photoelectron, and h is the energy of the incident photon. Although the X-ray may penetrate deeply into the sample to produce photoelectrons, most of these electrons lose energy in numerous inelastic collisions; only those atoms residing in the top few monolayers give rise to undistorted photoelectron spectra. The typical analysis depth in ESCA and in Ê , and they are truly surface analysis Auger electron spectroscopy is about 3±50 A methods. Auger electron spectroscopy (AES) In AES, an incident primary electron creates an excited ion near the surface which decays by the emission of a secondary Auger electron, whose kinetic energy is measured. As in XPS, the escaping Auger electron's kinetic energy limits the depth from which it can emerge, giving AES its high surface sensitivity and nanometer sampling depth. Auger images or maps can also be generated for specific elements with approximately 200 nm resolution. Auger

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finds its greatest strength in the analysis of inorganic materials not susceptible to electron-beam damage. Surface Raman Since its discovery in 1974, surface enhanced Raman scattering (SERS) has proved to be a very powerful and sensitive technique to investigate the vibrational properties of adsorbed molecules. It has been shown that the enhancement mechanisms can be of two kinds: a chemical phenomenon in which charge transfer between adsorbed molecules and the metallic substrate is involved and an electromagnetic enhancement which considers a localized surface plasmon. Yet, there is still some debate inside the SERS community as to whether one effect predominates or not (Grand et al., 2003). Since its original discovery, surface enhanced Raman scattering (SERS) has developed significantly towards detection of a single molecule. It was also more successful in biological applications than conventional Raman spectroscopy, especially in endoscopy and in-vivo diagnosis. Novel applications in biology include cancer gene detection, spectroscopy of living cells and single protein/ DNA detection. Progress in non-biological applications of SERS has been equally spectacular, with single molecule detection, spectroscopy of single dyes in large nanocrystals, or in carbon nanotubes (Etchegoin et al., 2003). Attenuated total reflectance (ATR) spectroscopy Before the advent of ESCA the only reliable method for polymer surface studies was infra-red spectroscopy (either attenuated total reflectance (ATR) or multiple internal reflectance (MIR) spectroscopy). This method requires fairly large samples with flat or easily deformable surfaces and typically gives information pertaining to surfaces and gives information penetrating to 1 m into the material. In recent years, Fourier transform infra-red (FTIR) has allowed analysis sensitivity to be improved, but even then FTIR can neither match the sensitivity of ESCA nor can it be as focused on the surface. In systems with variable composition within the range 0±1 m, the two methods could be applied to complement each other effectively. Infra-red microscopy Infra-red microscopy is a powerful technique that combines the image analysis capabilities of optical microscopy with the chemical analysis capabilities of infrared spectroscopy. The combination of these two techniques allows infra-red spectra to be obtained from microspectroscopic-sized samples. With the assistance of microscopy, samples as small as 0.01 g or even less (depending on the infra-red absorption characteristics of the components of interest) can be easily located and

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detected. Among its key roles are point identification or point mapping, using single element detectors to fingerprint defects and contaminants, analyze laminate structures, and to map chemical and physical property differences and gradients. These are critical to understanding product behavior, since chemical structure and conformation and physical property anisotropy at the microscopic level significantly influence the macroscopic performance of polymer products. A major limitation of FTIR microscopy has been the time needed using single element detectors to construct high spectral contrast maps of high spatial (lateral) resolution (~10 m) from large areas (e.g. >50 m  50 m). However, over the last few years, images based on mid-infra-red spectral differences and changes generated using FTIR microscopy instrumentation fitted with focal plane array (FPA) detectors have increasingly become available (Chalmers et al., 2002). In one application time-resolved FTIR measurements during isothermal crystallization of samples were carried out on a cast film. The crystallization of samples was traced by the C=O stretching band at 1722 cmÿ1 and the C±D stretching band at 2230 cmÿ1. Because of the large difference in molar absorbance coefficients, the C=O and C±D stretching bands were observed by separate measurements (He et al., 2002).

2.8.3 Energetics Surface energetics is very important in defining the biocompatibility of a polymer because this determines the type of molecules that do and do not adsorb on the surface. Tissue ingrowth, blood coagulation, cell damage to blood elements and immune signalling are all influenced by the energetics of the surface. Zeta potential A charged molecule in motion produces an electric field. The zeta potential is the electric potential of a charged particle at the plane of shear. The shear plane, or the plane of slip, is the distance from the surface to the distance in solution where the solvent molecules are not bound to the surface and are not moving as a unit with the particle. At this boundary, zeta potential can be determined. The surface potential is a very difficult parameter to characterize and it is far easier to determine zeta potential. It is measured at a shear plane near the particle surface and is therefore proportional to surface potential. It is affected by ionic strength; at high ionic strength, potential decreases much more sharply over the distance from the surface to the shear plane. This means a smaller zeta potential. It is therefore important to maintain a relatively constant ionic strength while characterizing the zeta potential of a dispersion as a function of pH. In the biomedical field, a large variety of applications are found. Gene therapy is one of them (Putnam et al., 2003). Complexation of plasmid DNA

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with polycations is a popular method by which to transfer therapeutic nucleic acid sequences to cells. One disadvantage of the approach is that the positive zeta potential of the complexes facilitates interaction with blood constituents, leading to serum protein adsorption and immune response. To circumvent this issue, investigators have developed polycations combined with polyethylene glycol (PEG) to create complexes with reduced protein adsorption potential. Contact angle The investigation of surface wettability by means of contact angle determination is of special interest in the characterization of the polymer surface. Contact angle may be geometrically defined as the angle formed by the intersection of two planes at a tangent to the liquid and solid surface at the perimeter of contact between the two phases and the third surrounding phase. Typically, the third phase will be air or vapor, although systems in which it is a second liquid essentially immiscible with the first are of great practical importance. If one considers the three-phase system where the liquid is designated as l, the surrounding gaseous medium as g, and the solid surface as s, then at equilibrium the contact angle  will be given by Young's equation as: gl
cos  ˆ sg ÿ sl where gl , sl and sg are the interfacial tensions at the respective interfaces. Contact angle measurements are carried out in various ways of widely differing sensitivity. Typically they are made with a goniometer and a syringe with a flat-tipped needle that is used to apply the solvent (generally double or triple distilled water) droplet on the surface. Advancing contact angles are recorded while fluid is added to the drop already on the surface. The values are generally reported as an average of five to ten measurements made on different areas of the sample surface. In culturing cells on biomaterial surfaces, contact angle is an important parameter that guides surface property modification attempts. For example, when oxygen plasma treated polyethylene surfaces were tested for tissue engineering purposes, the adherence of PC-12 cells were found to be higher as the water contact angle was lowered (higher surface wettability). However, this adhesion started to decrease as the contact angle was further decreased. Thus a bell-shaped curve for cell adhesion was obtained with the optimum water contact angle of 55 degrees (Lee et al., 2003).

2.9

Surface properties and biomaterials applications

Biomaterials are designed to augment, support or completely take over the functions of natural tissue or organs of human beings and are, therefore, designed to be biocompatible (non-toxic, non-immunogenic, non-inflammatory,

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non-carcinogenic) and to carry properties that match those of the tissue they are replacing. When functioning in the body the tissue-material interface is critical because this is where all the initial interactions such as contact with cells, proteins, and other biological entities takes place. The bulk of the biomaterial needs to satisfy the mechanical and physical requirements of the intended application, while an inert surface has to prevent immune responses, irritation or clotting. Biomaterial surfaces are, therefore, generally treated to improve their in-vivo responses. A typical example is the pyrolitic carbon-coated heart valve where the valve made of a metal or graphite is rendered hemocompatible and impact resistant by coating with a carbon layer. Heparin treated polymeric blood vessels and silicone sheathed sensor leads are all in this category. To improve blood compatibility it is a common approach to increase the hydrophilicity of the surface of the polymeric biomaterial and to achieve this surface modification by use of plasma treatment and subsequent grafting with a hydrophilic polymer. The optimum properties of a biomaterial can still not be elucidated, and a rule of thumb does not exist. The design of the surface of a biomaterial remains by far the most important and difficult task that a biomaterial scientist has to tackle.

2.10 Conclusion Surfaces of materials used in the biomedical field have always been important due to their proximity to the tissues. This is the first plane of contact between them and the nature of this contact is the main factor in determining their biocompatibility. Among the hottest current research trends today is nanobiotechnology where physical and chemical modification of biomaterial surfaces, their characterization and understanding of the nature of materialtissue interactions are of prime importance in today's biomaterials sciences. As the technology and our understanding of these increase, so will our ability to create better biomaterials. We will thus achieve higher quality of life.

2.11 References Ang TH, Sultana FSA, Hutmacher DW, Wong YS, Fuh JYH, Mob XM, Loh HT, Burdet E, Teoh SH (2002), `Fabrication of 3D chitosan±hydroxyapatite scaffolds using a robotic dispensing system', Materials Science and Engineering C, 20, 35±42. Barbu E, Pullin RA, Graham P, Eaton P, Ewen RJ, Smart JD, Nevell TG, Tsibouklis J (2002), `Poly(di-1H,1H,2H,2H-perfluoroalkyltaconate) films: surface organization phenomena, surface energy determinations and force of adhesion measurements', Polymer, 43, 1727±1734. Begovac PC, Thomson RC, Fisher JL, Hughson A, Gallhagen A (2003), `Improvements in GORE-TEX vascular graft performance by Carmeda BioActive surface heparin immobilization', Eur J Vasc Endovasc Surg, 25(5), 432±437. Casey BG, Monaghan W, Wilkinson CDW (1997), `Embossing of Nanoscale Features and Environments', Microelectronic Engineering, 35, 393±396.

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Chalmers JM, Everall NJ, Schaeberle MD, Levin IW, Lewis EN, Kidder LH, Wilson J, Crocombeg R (2002), `FT-IR imaging of polymers: an industrial appraisal', Vibrational Spectroscopy, 30(1), 43±52. Clough RL (2001), `High-energy radiation and polymers: A review of commercial processes and emerging applications', Nuclear Instruments and Methods in Physics Research, B, 185, 8±33. Davies MC, Shakesheff KM, Shard AG, Domb A, Roberts JC, Tendler SJB, Williams PM (1996), `Surface analysis of biodegradable polymer blends of poly(sebacicacid anhydride) and poly(DL-lactic acid)', Macromolecules, 29, 2205±2212. Etchegoin P, Maher R.C, Cohen L.F, Hartigan H, Brown RJC, Milton MJT, Gallop JC (2003), `New limits in ultrasensitive trace detection by surface enhanced Raman scattering (SERS)', Chemical Physics Letters, 375(1±2), 84±90. Fujita M, Mizutanic W, Gadd M, Shigekawab H, Tokumotoc H (2002), `Patterning DNA on mm scale on mica', Ultramicroscopy, 91, 281±285. Grand J, Kostcheev S, Bijeon J-L, de la Chapelle ML, Adam P-M, Rumyantseva A, LeÂrondel G, Royer P (2003), `Optimization of SERS-active substrates for near-field Raman spectroscopy', Synthetic Metals, 139(3), 621±624. Grobe GL, Valint PL, Ammon DM (1996), `Surface chemical structure for soft contact lenses as a function of polymer processing', J. Biomedical Materials Research, 32, 45±54. Harron HR, Pritchard RG, Cope BC, Goddard DT (1996), `An atomic force microscope (AFM) and tapping mode AFM study of the solvent induced crystallization of polycarbonate thin films', J. Polymer Science (Pt B): Polymer Physics, 34, 173±180. Hasirci V, Lewandrowski K, Bondre SP, Gresser JD, Trantolo DJ, Wise DL (2000), `High strength bioresorbable bone plates: preparation, mechanical properties and in vitro analysis', Bio-medical Materials and Engineering, 10(1), 19±29. Hasirci V, Litman AE, Trantolo DJ, Gresser JD, Wise DL, Margolis HC (2002), `Investigation of the interpenetrating network structure of a molecularly reinforced biodegradable implant', J. Materials Science, Materials in Medicine, 13, 159±167. Hasirci V, Bonney I, Goudas LC, Shuster L, Carr DB, Wise DL (2003), `Antihyperalgesic effect of simultaneously released hydromorphone and bupivacaine from polymer fibers in the rat chronic constriction injury model', Life Sciences, 73(26), 3323±3337. He W, Shanks R, Amarasinghe G (2002), `Analysis of additives in polymers by thin-layer chromatography coupled with Fourier transform-infrared microscopy', Vibrational Spectroscopy, 30(2),147±156. Jandt KD, McMaster TJ, Miles MJ, Petermann J (1993), `Scanning force microscopy of melt crystallized, metal evaporated poly(butene-1) ultrathin films', Macromolecules, 26, 6552±6556. Jandt KD, Buhk M, Miles MJ, Petermann J (1994), `Shish-kebab crystals in polyethylene investigated by scanning force microscopy', Polymer, 35(11), 2458±2462. Jeong B, Gutowska A (2002), `Lessons from nature: stimuli responsive polymers and their biomedical applications', Trends in Biotechnology, 20(7), 305±311. Jeong B, Kim, SW, Bae YH (2002), `Thermosensitive sol±gel reversible hydrogels', Advanced Drug Delivery Reviews, 54, 37±51. Kawai,T, Saito K, Lee W (2003), `Protein binding to polymer brush, based on ionexchange, hydrophobic, and affinity interactions', Journal of Chromatography B, Analyt Technol Biomed Life Sci, 790(1±2), 131±142. Klee D, HoÈcker H (2000), `Polymers for Biomedical Applications: Improvement of the

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Interface Compatibility', in Advances in Polymer Science, Springer-Verlag, Heidelberg, Vol. 149, 1±57. Lee SJ, Khang G, Lee YM, Lee HB (2003), `The effect of surface wettability on induction and growth of neurites from the PC-12 cell on a polymer surface', Journal of Colloid and Interface Science, 259, 228±235. Lustiger A, Lotz B, Duff TS (1989), `The morphology of the spherulitic surface in polyethylene', J. Polymer Science (Pt B): Polymer Physics, 27, 561±579. Meyer-Plath AA, Schroder K, Finke B, Ohl A (2003), `Current trends in biomaterial surface functionalization±nitrogen-containing plasma assisted processes with enhanced selectivity', Vacuum, 71, 391±406. Michanetzis GP, Katsala N, Missirlis YF (2003), `Comparison of haemocompatibility improvement of four polymeric biomaterials by two heparinization techniques', Biomaterials, 24(4), 677±688. Nam YS, Kang HS, Park JY, Park TG, Han S-H, Changa I-S (2003), `New micelle-like polymer aggregates made from PEI±PLGA diblock copolymers: micellar characteristics and cellular uptake', Biomaterials, 24(12), 2053±2059. Nirmala RJ, Jayakrishnan A (2003), `Surface thiocyanation of plasticized poly(vinyl chloride) and its effect on bacterial adhesion', Biomaterials, 24, 2205±2212. Pede D, Serra G, De Rossi D (1996), `Microfabrication of conducting polymer devices by ink-jet stereolithography', Materials Science and Engineering, C5, 289±291. Putnam D, Zelikin AN, Izumrudov VA, Langer R (2003), `Polyhistidine ± PEG:DNA nanocomposites for gene delivery', Biomaterials, 24(24), 4425±4433. Riccardia C, Barnia R, Sellib E, Mazzoneb G, Massafrac MR, Marcandallic B, Polettid G (2003), `Surface modification of poly(ethylene terephthalate) fibers induced by radio frequency air plasma treatment', Applied Surface Science, 211, 386±397. Sanchez JL, Guy G, van Kan JA, Osipowicz T, Watt F (1999), `Proton micromachining of substrate scaffolds for cellular and tissue engineering', Nuclear Instruments and Methods in Physics Research B, 158 185±189. Smith JR, Campbell SA, Mills GA (1997), `Probing atoms', Educ. Chem, 34(4), 107±111. Stachurski ZH (2003), `Strength and deformation of rigid polymers: structure and topology in amorphous polymers', Polymer, 44, 6059±6066. Tonge SR, Tighe BJ (2001), `Responsive hydrophobically associating polymers: a review of structure and properties', Advanced Drug Delivery Reviews, 53, 109±122. Toorisaka E, Uezua K, Goto M, Furusaki S (2003), `A molecularly imprinted polymer that shows enzymatic activity', Biochemical Engineering Journal, 14, 85±91. Tyan YC, Liao JD, Wu YT, Klauser R (2002), `Anticoagulant activity of immobilized heparin on the polypropylene nonwoven fabric surface depending upon the pH of processing environment', J Biomater Appl. 17(2), 153±178. Williams DF, The Williams Dictionary of Biomaterials, Liverpool University Press, 1999. Yamasaki A, Imamura Y, Kurita K, Iwasaki Y, Nakabayashi N, Ishihara K (2003), `Surface mobility of polymers having phosphorylcholine groups connected with various bridging units and their protein adsorption-resistance properties', Colloids and Surfaces B: Biointerfaces, 28, 53±62. Zeina I, Hutmacher DW, Tanc KC, Teoh SH (2002), `Fused deposition modeling of novel scaffold architectures for tissue engineering applications', Biomaterials, 23, 1169±1185. Zhu A, Zhang M, Wu J, Shen J (2002), `Covalent immobilization of chitosan/heparin complex with a photosensitive hetero-bifunctional crosslinking reagent on PLA surface', Biomaterials, 23(23), 4657±4665.

3

Organic thin film architectures: fabrication and properties M C P E T T Y , University of Durham, UK

3.1

Introduction

Molecular electronics is a fast moving and interdisciplinary subject that exploits the electronic and optoelectronic properties of organic and biological materials (Petty et al., 1995; Richardson, 2000; Tour, 2003; Maruccio et al., 2004). The areas of application and potential application are varied, ranging from chemical and biochemical sensors to plastic light-emitting displays. Molecular electronics also offers considerable scope for molecular nanotechnology, e.g., the development of electronic switching and memory structures operating, and addressable, at the molecular level. In most of these examples, the organic materials are required in the form of thin films (1 nm to 10 m). This presents a considerable challenge to materials scientists, as organic compounds, in their bulk form, can be fragile and difficult to handle. In this chapter, an overview of the more popular methods that may be used to fabricate thin layers of organic compounds will first be described. A distinction will be drawn between established deposition technologies, in many cases developed for use with inorganic materials, and those methods that actually allow molecular-scale architectures to be built-up on solid supports. Each deposition method differs in complexity and may be more suited to provide films in a particular thickness range. Specific types of organic compound are necessary for certain processes, e.g., self-assembly exploits the attraction between certain chemical groups. Some methods are `wet' (spinning) while others are inherently `dry' (plasma deposition). There are also implications for the degree of order and contamination levels in the deposited film. Following the introduction to organic film technologies, a brief review of some of the more powerful analytical techniques that can be used to reveal the structure and degree of molecular organization in the thin layers will be presented.

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3.2

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Established deposition methods

3.2.1 Physical vapour deposition Solid materials vaporize when heated to sufficiently high temperatures, this process may proceed through the liquid phase. A thin film is then obtained by the condensation of the vapour onto a colder substrate (Maissel and Glang, 1970). This method has been used extensively to deposit films of inorganic materials, such as metals and their alloys. However, the technique is now being used for the formation of layers of low molecular weight organic compounds. Because of collisions with ambient gas atoms, a fraction of the vapour atoms will be scattered. For a straight line path between the evaporating material (source) and the substrate, it is necessary to use low pressures (< 10ÿ4 mbar) where the mean free path of the gas atoms is much greater than the sourcesubstrate distance. This allows the use of a shadow mask immediately in front of the substrate to define patterns. The low pressure also prevents contamination of the source material (e.g. by oxidation). Figure 3.1 shows a schematic diagram of a typical evaporation system. The system chamber, which can be made from glass or metal, is evacuated to a pressure of 10ÿ4±10ÿ6 mbar, normally with two types of vacuum pump, a rotary and diffusion pump, operating in series. The first step in physical vapour deposition requires the transformation of the condensed phase, solid or liquid, into the gaseous state. This conversion of thermal to mechanical energy can be achieved by a variety of methods. Resistive heating has been used to deposit fluorescent dyes, charge-transfer salts and large

3.1 Schematic diagram of a typical vacuum evaporation system for physical vapour deposition.

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macromolecules, such as the phthalocyanines (Petty, 2000; Park et al., 2004). Typical evaporation rates are 1±10 nm minÿ1. Other techniques include arc evaporation, RF heating or heating by electron bombardment. Deposition of polymer films by laser ablation is an area that offers some promise (Chrisey and Hubler, 1994). Laser pulsed methods have been used successfully for polyethylene, polycarbonate, polyimide, polymethylmethacrylate and, more recently, for the deposition of compounds for organic light-emitting displays and transistors (Salih et al., 1997; Hong et al., 2002; Blanchet et al., 2003). Materials that dissociate in the vapour phase may provide solid films with a stoichiometry that differs from that of the source. Therefore special techniques have been devised. One approach is to use the method of `flash' evaporation. It is also possible to evaporate from two, or more, sources and to control the flux from each to obtain a vapour with the required composition. This has been used effectively to deposit thin films of doped organic charge-transfer salts (Breen et al., 1993): one source is the charge-transfer salt, e.g., tetrathiafulvalene (TTF), while the other is the dopant, e.g., iodine. Laser co-ablation techniques can be used to deposit films of metal-polymer composites (Chrisey and Hubler, 1994). Molecular beam epitaxy (MBE) is a similar, but more expensive, variation of vacuum evaporation (Hara and Sasabe, 1995). However, an ultra-high vacuum (<10ÿ9 mbar) is required to eliminate the scattering by residual gas molecules. The technique consists of directing controlled `beams' of the required molecules towards a heated substrate. Multiple sources, Knudsen cells, can be shuttered and used to create a superlattice structure on the substrate and control the molecular composition, orientation and packing in two dimensions. Some of the first trial observations of real-time epitaxial growth in organic molecular systems have involved a study of phthalocyanine monolayers on cleaved surfaces of MoS2 and on highly orientated pyrolyzed graphite and alkali halides. In-situ reflection high energy diffraction (RHEED) can be used to monitor the actual film growth (Hara and Sasabe, 1995). Highly ordered thin film crystals of perylene derivatives and electrically conductive oligomers for the fabrication of organic field effect transistors have also been produced by MBE (Dimitrakopoulos et al., 1998; Sazaki et al., 2004). A further physical vapour deposition technique, sputtering, is based on the momentum exchange of accelerated ions incident on a target of source material (Maissel and Glang, 1970). A source of ions (e.g. Ar or Xe) is provided by a glow discharge created by an electric field between two electrodes in a gas at low pressure. The material to be sputtered, the target, is the cathode. The vacuum chamber is evacuated and then filled with the inert gas at a low pressure. A potential difference of several kilovolts is applied to the electrodes which causes the gas to become ionized and form a plasma. The positive gas ions are then accelerated by the electric field so that they arrive at the cathode with considerable energy and sputter the target atoms. Some secondary electrons are also produced at the cathode and these accelerate towards the anode and serve to

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maintain the plasma. There are several variations on the basic sputtering arrangement described above. For example, RF (radio frequency) sputtering can be used to sputter insulating source materials. The principal advantage of sputtering is that almost any material can be deposited. Since no heating is required, materials that are difficult to melt (and therefore to evaporate) are easily sputtered, as are compounds that would dissociate in an evaporation source. The method has been used for some organic polymers such as polyimide (Kinbara et al., 2003) and for producing fluorescent coatings (Maggioni et al., 2003). However, a relatively large amount of the material is needed as a target. This is not always practical for the new organic compounds, which may be available only in very small quantities.

3.2.2 Spin-coating Spin-coating is exploited by the microelectronics industry for depositing layers of photoresist films, generally polyimides, onto silicon wafers (Brodie and Muray, 1992). The various steps involved are illustrated in Fig. 3.2. A quantity of a polymer solution is first placed on the semiconductor wafer, which is then rotated at a fixed speed of several thousand rpm (or the solution can be applied while the wafer is slowly rotating). The resist solution flows radially outwards, reducing the fluid layer thickness. Evaporation of the solvent results in a film of uniform thickness. The resist viscosity (dependent on the concentration of the starting solution) and final film speed are both important process parameters that influence the final film thickness. An increase in angular velocity decreases the film thickness; an inverse power-law relationship usually holds for the thickness dependence on the final spin speed. For a given speed, the film thickness decreases rapidly at first, but then slows considerably at longer times. Although spin-coating is expected to produce films in which individual molecules are relatively disordered, this is not always the case. For instance,

3.2 Schematic representation of the spin-coating technique.

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spin-coated phthalocyanine layers have revealed crystalline order (Critchley et al., 1992) which may result from the centrifugal forces acting upon the individual molecules during spinning. In other cases, order can be achieved by a post-deposition treatment. Application of heat and an electric field normal to film plane (the process of `poling') can be used to align the C-F dipoles in polyvinylidene fluoride films. The result is a polar film possessing piezoelectric and pyroelectric behaviour. A similar method can be used to induce secondorder nonlinear optical behaviour (e.g. second-harmonic generation) in films of appropriate molecules (Petty et al., 1995).

3.2.3 Other techniques Chemical vapour deposition (CVD) is a process in which one or more gaseous species react on a solid surface and one or more of the reaction products is a solid phase material (Maissel and Glang, 1970; Sherman, 1987). The CVD process is based on the decomposition and/or radical generation of chemical species by stimulating vapour with heat, plasma (discharge) or light (laser). The method is used in the microelectronics industry for the fabrication of inorganic semiconducting and insulating films. In plasma-enhanced CVD (PECVD) glow discharge plasmas are sustained within chambers where simultaneous CVD reactions occur (Yasuda, 1985; Lucovsky et al., 1990). The discharge is normally excited by an RF field. One example is the formation of polystyrene films as the dielectric in a nuclear battery (Brodie and Muray, 1992; Lucovsky et al., 1990). The surface of the polymer may be simultaneously removed by reaction with the residual gases (ablation); the balance between ablation and polymerization may be controlled by adding gases that promote one or other effect. For example, CF4 is an effective plasma etch material under normal discharge conditions; however, if a small amount of hydrogen is added, a polymer is deposited. Laser, or more generally, optical chemical processing involves the use of monochromatic photons to enhance and control reactions at substrates (Eden, 1992). Electrochemical deposition, or electroplating, has been known for at least one hundred years. Approximately half of the 70 or so metals can be electrodeposited, either singly or as alloys. The equipment required consists of an anode and a cathode immersed in a suitable electrolyte. Metal is deposited onto the cathode and the relationship between the weight of the material deposited and the various parameters can be expressed by the first and second laws of electrolysis. Anodization is a particular kind of electrolysis. Many polymeric films can be prepared by the anodic oxidation of suitable monomer species, such as pyrrole, thiophene and aniline (Xia et al., 2001; Grigore and Petty, 2003; Alumaa et al., 2004). The method has also been used to prepare thin layers of biopolymers (Kastantin et al., 2003).

Organic thin film architectures: fabrication and properties

3.3

65

Molecular architectures

3.3.1 Langmuir-Blodgett technique A method that allows the manipulation of materials on the nanometre scale is the Langmuir-Blodgett (LB) technique (Petty, 1996). Langmuir-Blodgett films are prepared by first depositing a small quantity of an amphiphilic compound (i.e. one incorporating both polar and nonpolar groups ± the classical materials being long-chain fatty acids) dissolved in a volatile solvent, onto the surface of purified water, the subphase. When the solvent has evaporated, the organic molecules may be compressed to form a floating two-dimensional solid. Monolayer films can exhibit a multiplicity of phases during this compression; these are similar to the mesophases shown by liquid crystals. The surface pressure  of the floating organic film is defined as the reduction of the subphase surface tension by the film, i.e.  ˆ o ÿ

3:1

where o is the surface tension of the pure subphase and is the surface tension of the film-covered surface. In the case of a water subphase (nearly almost all the work on LB films has involved a water subphase) values of  of the order of mN mÿ1 are generally encountered; the maximum value of  is 72.8 mN mÿ1 at 20 ëC, the surface tension of pure water. Figure 3.3 shows a surface pressure versus area isotherm (i.e. measurement at constant temperature) for a hypothetical long-chain organic material. This generic diagram is not meant to represent that observed for a particular

3.3 Surface pressure versus area per molecule isotherm for a long-chain organic compound. The surface pressure and area are in arbitrary units (a.u.).

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substance, but shows most of the features observed for long-chain compounds (Petty, 1996). In the `gaseous' state (G in Fig. 3.3) the molecules are far enough apart on the water surface that they exert little force on one another. As the surface area of the monolayer is reduced, the hydrocarbon chains will begin to interact. The `liquid' state that is formed is generally called the expanded monolayer phase (E). The hydrocarbon chains of the molecules in such a film are in a random, rather than a regular orientation, with their polar groups in contact with the subphase. As the molecular area is progressively reduced, condensed (C) phases may appear; there may be more than one of these. In the condensed monolayer states, the molecules are closely packed and are orientated with the hydrocarbon chains pointing away from the water surface. The limiting area per molecule in such a state will be similar to the cross-sectional area of the hydrocarbon chain, i.e. approximately 0.19 nm2 moleculeÿ1. If the surface pressure of the monolayer is held constant in one or more of the condensed phases, then the film may be transferred from the water surface onto a suitable solid substrate simply by raising and lowering the latter through the monolayer/air interface. Monolayer transfer speeds are usually of the order of a few mm per second. Several deposition modes are possible depending on the interactions between the polar and nonpolar parts of the molecules, and the nature of the bond between the first layer and the substrate surface. In the commonest form of LB film deposition, known as Y-type deposition, a monolayer is deposited on each traversal of the monolayer/air interface. Instances in which the floating monolayer is only transferred to the substrate as it is being inserted into the subphase, or only as it is being removed are often observed. These deposition modes are called X-type (monolayer transfer on the downstroke only) and Z-type (transfer on the upstroke only). Film transfer is characterized by the measurement of a deposition, or transfer, ratio, . This is the decrease in the area occupied by the monolayer (held at constant pressure) on the water surface divided by the coated area of the solid substrate, i.e., ˆ

AL AS

3:2

where AL is the decrease in the area occupied by the monolayer on the water surface and AS is the coated area of the solid substrate. Transfer ratios significantly outside the range 0.95 to 1.05 suggest poor homogeneity. A schematic diagram of one possible experimental arrangement for the deposition of LB films is shown in Fig. 3.4. The container that holds the water is termed a Langmuir trough. A working area can be defined by a PTFE-coated glass fibre barrier that can be moved using a low-geared electric motor. The barrier motor is coupled to a sensitive electronic balance which continuously monitors, via a sensing plate, the surface pressure of the monolayer. Using a feedback arrangement, this pressure can be maintained at a predetermined value. The physical dimensions of the Langmuir trough are not critical and are governed by

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3.4 Schematic diagram of LB trough and associated monolayer compression system for multilayer deposition.

the size of the substrate used. A dipping well can be used to accommodate large substrates without increasing the volume of water in the trough. Cleanliness is essential to produce good quality films. To minimize contamination, the system is usually housed in a glove box or in a microelectronics clean room. It is also possible to construct LB films using more than one type of molecular film. In the simplest case, condensed floating monolayers of two different amphiphilic materials are confined (using mechanical barriers) to different regions of the water surface. By lowering the solid substrate through the first layer of, say material A, and raising it up through the other, material B, alternate-layers of structure ABABAB. . . may be built up. This permits the fabrication of organic superlattices with precisely defined symmetry properties.

3.3.2 Self-assembly Self-assembly is a much simpler process than that of LB deposition. Monomolecular layers are formed by the immersion of an appropriate substrate into a solution of the organic material (Fig. 3.5). The best known examples of self-assembled systems are organosilicon on hydroxylated surfaces (SiO2, Al2O3, glass, etc.) and alkanethiols on gold, silver and copper (Ulman, 1991; Tredgold, 1994). However, other combinations include: dialkyl sulphides on gold; dialkyl disulphides on gold; alcohols and amines on platinum; and carboxylic acids on aluminium oxide and silver. The self-assembly process is driven by the interactions between the head group of the self-assembling molecule and the substrate, resulting in a strong chemical bond between the head

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3.5 Self-assembled monolayer of an alkanethiol on a gold-coated substrate.

group and a specific surface site, e.g., a covalent Si-O bond for alkyltrichlorosilanes on hydroxylated surfaces. The self-assembly process is usually restricted to the deposition of a single molecular layer on a solid substrate. However, chemical means can be exploited to build up multilayer organic films. A method pioneered by Sagiv is based on the successive absorption and reaction of appropriate molecules (Netzer and Sagiv, 1983; Netzer et al., 1983). The headgroups react with the substrate to give a permanent chemical attachment and each subsequent layer is chemically attached to the one before in a very similar way to that used in systems for supported synthesis of proteins.

3.3.3 Electrostatic layer-by-layer deposition Another technique for building up thin films of organic molecules is driven by the ionic attraction between opposite charges in two different polyelectrolytes. Figure 3.6 shows a schematic diagram of the processes involved in this layer-by-

3.6 Sequence of immersion of a substrate in polyanion and polycation solutions for electrostatic layer-by-layer deposition.

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3.7 Schematic diagrams of expected molecular organization in (a) a LangmuirBlodgett film and (b) a film deposited by the layer-by-layer electrostatic technique.

layer (LbL) assembly technique (Decher and Schlenoff, 2003). A solid substrate with a positively charged planar surface is immersed in a solution containing an anionic polyelectrolyte and a monolayer of polyanion is adsorbed. Since the adsorption is carried out at relatively high concentrations of the polyelectrolyte, most of the ionic groups remain exposed to the interface with the solution and thus the surface charge is reversed. After rinsing in pure water, the substrate is immersed in a solution containing the cationic polyelectrolyte. Again, a monolayer is adsorbed but now the original surface charge is restored, thus resulting in the formation of multilayer assemblies of both polymers. It is possible to use a sensitive optical technique, such as surface plasmon resonance to monitor, in situ, the growth of such electrostatically assembled films (Pearson et al., 2001). The layer-by-layer method has been used to build up layers of conductive polymers, e.g., partially doped polyaniline and a polystyrene polyanion (Cheung et al., 1997). Biocompatible surfaces consisting of alternate layers of charged polysaccharides and oppositely charged synthetic polymers can also be deposited in this way (Lvov et al., 1998). A related, but alternative, approach uses layer-by-layer adsorption driven by hydrogen bonding interactions (Stockton and Rubner, 1997). Figure 3.7 contrasts the molecular organization expected in an LB multilayer with that in electrostatically deposited layer-by-layer films. For the latter case, the polyelectrolyte chains within each layer will become entangled, and may even penetrate into the layers above and below, leading to a less ordered film than that produced by LB deposition.

3.3.4 Dip-pen nanolithography Scanning microscopy methods offer a powerful means of manipulating molecules. For instance, careful control of an atomic force microscope (AFM) tip can allow patterns to be drawn in an organic film (Chi et al., 1992). Such

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3.8 Dip-pen patterning technique showing the transfer of an `ink' onto `paper' using the tip of an atomic force microscope (Piner et al., 1999).

techniques can also be used to reposition molecules, such as the fullerene C60, on surfaces and to break up an individual molecule (Gimzewski, 1998). A further approach that has been developed at Northwestern University is called Dip-Pen Nanolithography (DPN) (Piner et al., 1999). This technique, illustrated in Fig. 3.8, is able to deliver organic molecules in a positive printing mode. An AFM tip is used to `write' alkanethiols on a thin gold film in a manner analogous to that of a fountain pen. Molecules flow from the AFM tip to the solid substrate (`paper') via capillary transport, making DPN a potentially useful tool for assembling nanoscale devices. The chemisorption of the `ink' is the driving force that moves the ink from the AFM tip through the water to the substrate as the tip is scanned across this surface. Adjusting the scan rate and relative humidity can control line widths. Recent developments of DPN have included an overwriting capability that allows one nanostructure to be generated and the areas surrounding it to be filled with a second type of ink (Hong et al., 1999). Perhaps the greatest limitation in using scanning probe methodologies for ultra-high resolution nanolithography over large areas derives from the serial nature of most techniques. However, an eight-pen nanoplotter capable of doing parallel lithography has been reported (Hong and Mirkin, 2000). Parallel dip-pen nanolithography with arrays of individually addressable cantilevers has also been developed (Bullen et al., 2004). The DPN method has also been used to deposit magnetic nanostructures (Liu et al., 2002), arrays of protein molecules (Lee et al., 2002) and metal nanocrystals (Thomas et al., 2004).

3.3.5 Ink-jet printing The need to combine large area coatings with device patterning has resulted in the development of other direct-write fabrication methods, such as ink-jet printing. This particular method is attracting the interest of many workers worldwide and is currently a rapidly moving area of research and development, producing many papers and patents (Speakman et al., 2001; PercËin and Khuri-

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Yakub, 2002; Sirringhaus and Shimoda, 2003). Although ink-jet printhead droplet ejection can be achieved with thermal (bubble-jet) and piezoelectric modes of operation, the majority of published literature on ink-jet printing as a tool for manufacturing organic devices has been the result of using piezoelectric actuated printers. Piezoelectric printhead technology is favoured primarily because it applies no thermal load to the organic inks and is compatible with the printing of digital images. The combination of solution-processable emissive polymers with ink-jet printing offers some promise in the development of lowcost, high-resolution displays (Rees et al., 2002). The technique has also been applied to the manufacture of all-polymer transistor circuits (Sirringhaus et al., 2001; Paul et al., 2003).

3.4

Molecular organization in thin films

The need to understand the structure of organic thin films is paramount if these thin layers are to be exploited commercially. Traditional techniques such as optical and electron microscopies allow the morphology of the layers to be explored, while spectroscopic examination using electromagnetic radiation, electrons or ions can reveal the chemical nature of the films (Ulman, 1994; Tredgold, 1994). In the case of LB or self-assembled layers, special methods, such as X-ray and electron diffraction can show details of the molecular organization in both the out-of-plane and in-plane directions. Generally, two levels of molecular packing must be considered for long-chain molecules in layered structures. The organization of the individual molecules in the multilayer will define the main crystallographic cell while the packing of the C2H4 repeat units in the long hydrocarbon chains defines the subcell (Petty, 1996). The following section provides an overview of the methods that are commonly used to study organic thin films, with an emphasis on those techniques that can reveal useful information about molecular scale architectures.

3.4.1 X-ray, neutron and electron diffraction X-ray diffraction was one of the original methods used to study the structure of fatty acid type LB films. When electromagnetic energy is incident on the surface of a material, some of it will be reflected specularly so that the angle of incidence is equal to the angle of reflection. For crystalline materials, constructive and destructive interference between the radiation reflected from successive crystal planes will occur when the wavelength of the incident radiation is of the same order as the lattice spacing. The condition for maxima in the reflected radiation is provided by Bragg's law n ˆ 2dhkl sin 

3:3

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where  is the angle of incidence, conventionally measured from the plane of reflection in X-ray crystallography, dhkl is the interplanar spacing and the integer n is known as the order of reflection. Figure 3.9 shows X-ray diffraction data from a 43-layer fatty acid salt film (Nicklow et al., 1981). As predicted, the first order Bragg reflection is observed for an angle of incidence close to one degree. In a Y-type LB film, the d001 spacing is equal to the distance between the polar planes, i.e., the thickness of two monomolecular layers. The monolayer thicknesses for fatty acid salt films, obtained from X-ray experiments, are in close agreement with the lengths of the molecules, implying that the hydrocarbon chains in transferred monolayers are orientated at right-angles to the substrate surface. For Y-type LB films of other amphiphilic materials (including simple fatty acids) the X-ray d-spacing is often less than twice the molecular length, suggesting some tilting or interdigitation (or both) of the molecules. X-rays are scattered by their interaction with atomic electrons and interference takes place between X-rays scattered from different parts of an atom. The scattering power or scattering factor decreases with increasing scattering angle, 2, resulting in a decrease in the intensity of the Bragg peaks. However, it

3.9 X-ray diffraction data from a 43-layer LB film of perdeuterated manganese octadecanoate on a silicon substrate. Experimental values are shown as points. The solid curve is based on calculation and is displaced from the data points. (Reprinted from Nicklow, R. M., Pomerantz, M. and Segmuller, A., Physical Review B, 23, 1081±7, 1981. Copyright 1981 by the American Physical Society.)

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is evident that the height of the third Bragg reflection in Fig. 3.9 is greater than that of the second. This is because the intensity of a particular Bragg peak is not only related to the scattering power of each atom in the lattice but also to the position of the atom in the unit cell. In summing the individual waves to give the resultant diffracted beam, both the amplitude and phase of each wave scattered by the individual atoms are important. The intensity of the scattered radiation Ihkl from a set of planes `hkl' may be written 2 Ihkl ˆ sLpFhkl

3:4

where s is a scale factor, L is the Lorentz (geometrical) correction, p is a polarization correction and Fhkl is called the structure factor. The structure factor of the unit cell depends upon the constituent atoms and their individual scattering factors. For a particular d-spacing, the only variable in the cell structure is the nature of the atom. Since Fhkl depends directly on the effective number of scattering electrons per atom, a large Fhkl is given by a high atomic number. Thus fatty acid salt films, containing heavy metals such as cadmium, barium or lead, are generally superior for X-ray diffraction experiments. If the structure factors for a complete set of X-ray reflections are known, the electron density at any position in the unit cell may be calculated. However, the computational problem is formidable. A physical criterion (e.g. a region, such as an alkyl chain, where a constant electron density is expected) is often used to evaluate the electron density profile of multilayer films perpendicular to the substrate. Interference of reflected X-rays from the upper and lower surfaces of a thin film gives rise to Kiessig fringes. These may be used to determine the overall thickness of the film. The use of a neutron beam produces interference effects in a similar way to an X-ray beam. The major difference between the two types of radiation lies in the factors governing the intensity of diffraction. For X-rays, this depends on the electron density variation across the layers. In the case of neutrons, it is the variation of nuclear scattering length density that determines the Bragg intensity. Because neutron scattering is a nuclear property, it may vary considerably from one element to the next and is different for different isotopes, e.g., hydrogen and deuterium. Neutron absorption is usually negligible and the interference effects that cause X-ray scattering to diminish with increasing angle are absent with neutrons and the scattering is isotropic. Electron diffraction provides structural information on multilayer films in much the same way as X-ray experiments. To obtain the structure normal to the film plane, the electron beam impinges at a grazing angle and the reflected beam is observed. For fatty acid type LB films, the diffraction experiments reveal the packing of the C2H4 subcells in the aliphatic chain. One advantage of using electrons as a probe is that the interaction of the electron beam with matter is much stronger than for X-rays. Both transmission and reflection diffraction techniques can be used. Figure 3.10 shows a transmission electron diffraction

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3.10 Transmission electron diffraction pattern for a 22-tricosenoic acid LB film. (After Peterson and Russell, 1984. Reproduced with permission from Taylor & Francis Ltd: http://www.tandf.co.uk/journals.)

pattern of a fatty acid LB film (Peterson and Russell, 1984). This particular diffraction pattern can be indexed as arising from an orthorhombic packing of the subcells. The measurement of reflection high energy electron diffraction (RHEED) patterns in the direction of film deposition and perpendicular to it can reveal the presence of in-plane anisotropy and can also provide the tilt elevation of molecules in the LB film. Figure 3.11 shows that the angle of inclination to the horizontal varies with the deposition surface pressure (Peterson et al., 1988; Barnes and Sambles, 1987). As the deposition pressure is increased, the molecular tilt elevation also increases (i.e. the molecules become more upright).

3.11 Molecular tilt (measured from substrate plane) versus deposition surface pressure for 22-tricosenoic acid. n deposition pH7; l pH3; t pH7 (Barnes and Sambles, 1987) (after Peterson et al., 1988, Copyright 1988, with permission from Elsevier).

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3.4.2 Infra-red spectroscopy Infra-red (IR) spectroscopy (wavelength in the range 1±100 m) probes the vibrational features of an organic molecule. A particular chemical bond must have a permanent dipole moment associated with it to interact with the infra-red radiation. The compression and extension of the bond can be likened to the behaviour of a spring. This analogy may be taken further by assuming that the bond, like a simple spring, obeys Hooke's law. For a diatomic molecule (e.g. HCl) the vibrational energy levels E are given by   1 E ˆ h n ‡ 3:5 2 where h is Planck's constant,  is the frequency of the radiation and the vibrational quantum number n ˆ 0; 1; 2; . . .. For IR spectroscopy, it is usual for the energy of the electromagnetic radiation to be quoted in terms of a wavenumber . This is the reciprocal of the wavelength, in centimetres. Although a normal mode of vibration involves movement of all of the atoms in a molecule, there are circumstances in which movement is almost localized in one part of the molecule. If the vibration involves the stretching or bending of a terminal XY group, where X is heavy compared to Y, e.g., an OH group in a fatty acid, the corresponding vibration wavenumbers are almost independent of the rest of the molecule to which XY is attached. A typical wavenumber for the XY stretching vibration may therefore be referred to. For example, the OH stretching frequency is normally in the region of 3600 cmÿ1. Many group vibrations occur in the region 1500 to 3700 cmÿ1; the stretching and bending vibrations of some well-known groups are listed in Table 3.1. The absorption or reflection intensities resulting from the interaction of IR radiation with monolayer samples are very low, a direct result of the relatively Table 3.1 Characteristic stretching and bending frequencies of molecular groups Group

Approximate wavenumber (cmÿ1)

±OH ±NH2 =CH2 ±CH3

3600 3400 3030 2960 (antisym. stretch) 2870 (sym. stretch) 1460 (antisym. bend) 1375 (sym. bend) 2920 (antisym. stretch) 2850 (sym. stretch) 1470 (bend) 2220 1750±1600 1650

±CH2± ±C=C± >C=O >C=C<

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3.12 Comparison of (a) simple transmission (b) ATR and (c) RAIRS sampling techniques for infra-red spectroscopy. The organic thin films are indicated by the shaded regions.

small numbers of molecules being sampled. In a transmission experiment with normal incidence of the IR beam, the electric field is orientated parallel to the layer plane, as shown in Fig. 3.12(a). In this geometry the projection of the transition moments on the layer plane is probed. An increase in surface sensitivity may be obtained by using a method based on attenuated total reflection (ATR) (Fig. 3.12(b)). In this technique, the sample is deposited onto either side of an IR-transmitting crystal (e.g. silicon or germanium). The radiation is incident at an angle greater than the critical angle and undergoes multiple reflections inside the crystal. On each reflection, the evanescent field of the IR beam penetrates the organic layer and may be absorbed by it. Both the transmission and ATR experiment may be performed using either polarized or unpolarized radiation. Reflection absorption infra-red spectroscopy, RAIRS, requires IR radiation to be incident at a grazing angle to a metal surface onto which the organic film has been deposited (Fig. 3.12(c)). The RAIRS experiment has the very useful ability to detect vibrations that possess a transition dipole with a large component perpendicular to the surface. The orientation of the molecules in organized thin films can be investigated by using different polarizations of incident radiation or by comparing transmission and reflection measurements. In the case of the ATR technique, unpolarized IR radiation will couple to all the molecular dipoles a thin film. In contrast, RAIRS preferentially selects those transition dipoles with large components perpendicular to the film plane. Figure 3.13 shows the

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3.13 Infra-red spectra of 21 LB layers of cadmium docosanoate on a silicon ATR crystal. (a) ATR mode. (b) RAIRS mode. (Reprinted with permission from Davies and Yarwood, 1989, Copyright 1989 American Chemical Society.)

ATR and RAIRS spectra in the CH stretching region for a 21 layer fatty acid salt LB film (Davies and Yarwood, 1989). The bands that are observed are 2960 cmÿ1 (antisymmetric CH3 stretch), 2920 cmÿ1 (antisymmetric CH2 stretch), 2870 cmÿ1 (symmetric CH3 stretch) and 2850 cmÿ1 (symmetric CH2 stretch). The symmetric and antisymmetric CH2 stretches are both intense in the ATR spectrum. However, in the RAIRS mode, the CH3 bands increase in intensity. The transition dipole moments of the CH2 vibrations are perpendicular to the long axis of the fatty acid molecules, whereas the symmetric and antisymmetric CH3 stretches have components along the chain axis. Therefore, the experimental data shown in Fig. 3.13 are consistent with the long axes of the fatty acid molecules being almost perpendicular to the substrate plane.

3.4.3 Scanning probe microscopies Scanning tunnelling microscopy (STM) and atomic force microscopy (AFM) are both techniques that may be used to provide direct images of metal surfaces with nanometre resolution. Whereas STM records the overlap of the local electron density of states between tip and substrate or its modulation by adsorbate molecules in the gap, AFM measures the interatomic forces between a cantilevered spring tip and the sample surface. The image contrast in AFM is achieved by probing the elastic response of the molecules to the force exerted by the scanning tip. Figure 3.14 shows an example of an atomic force micrograph of a 12-layer neicosanoic acid LB film deposited onto single crystal silicon (Evenson et al., 1996). Lines of individual molecules are evident at the magnification shown.

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3.14 AFM image of a 12-layer n-eicosanoic acid LB film deposited onto silicon. (Reprinted with permission from Evenson et al., 1996, Copyright 1996 American Chemical Society.)

3.5

Future trends

The past three decades have witnessed a marked increase in academic and commercial interest in organized thin organic films. Much of the original work was devoted to understanding the formation and structure of Langmuir-Blodgett films. As applications for these ultra-thin layers have become a possibility, other techniques such as chemical and electrostatic self-assembly have been introduced. All these methods can provide intriguing molecular architectures which can form the basis of fundamental experiments in the research laboratory. For molecular nanotechnology to become a reality, there is still much to be discovered about the nature of the organic compounds that can be used in these techniques and the links between the film structure and its properties. This may require the development or refinement of characterization tools suited to the study of molecular assemblies. In the shorter term, the more established techniques of spin-coating and thermal evaporation may provide cost-effective solutions in the commercial arena for application areas in which a high degree of molecular order is not essential. The long-term stability of organic materials is still a problem for the commercialization of electronic devices. For example, organic light emitting displays require an operating lifetime of about 50,000 hours. The organic thin films that are used in these (and other) devices are very sensitive to contamination, oxidation and humidity. To increase the lifetime of such devices, various methods have been devised for their encapsulation. Hermetic encapsulation using a glass or metal lid has been extensively used to protect the organic films from oxygen and water vapour. However, these approaches are

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expensive. Other methods based on dry and wet processes are currently being explored. The stability issues associated with organic materials should not, however, deter scientists and technologists in their endeavours to develop products. The commercialization of liquid crystal displays demonstrates clearly that quite sophisticated devices incorporating organic thin films can be marketed. There are, of course, numerous examples of the successful exploitation of organized molecular assemblies in the natural world!

3.6

Further information

The work described in this chapter is highly interdisciplinary and, therefore, further information is to be found in a variety of physics, chemistry and engineering databases. Biologists and biochemists are also taking a keen interest in organized organic films as some of these can possess similar structures to cell membranes found in the natural world. Existing books on organized molecular organic films include those of Ulman (1991), Petty (1996) and Decher and Schlenoff (2003) while the properties and possible applications of these thin layers are covered by Tredgold (1994), Petty et al. (1995) and Richardson (2000). Developments in the fabrication, investigation and applications of organic thin films are the concerns of the International Conference on Organized Molecular Films (LB series) and the European Conference on Organized Films (ECOF series). Both take place every two or three years and their proceedings are published in journals such as Thin Solid Films and Materials Science and Engineering C. In addition, many of the activities are covered in the nanotechnology literature. A good starting source of information is the website of the UK-based Institute of Nanotechnolgy (http://www.nano.org.uk/).

3.7

References

Alumaa A, Hallik A, Maeorg U, Sammelselg V and Tamm J (2004) Potentiometric properties of polypyrrole bilayers, Electrochimica Acta, 49, 1767±1774. Barnes W L and Sambles J R (1987) Surface pressure effects on Langmuir-Blodgett multilayers of 22-tricosenoic acid, Surface Sci., 187, 144±152. Blanchet G B, Fincher C R and Malajovich I (2003) Laser evaporation and the production of pentacene films, J. Appl. Phys., 94, 6181±6184. Breen J J, Tolman J S and Flynn G W (1993) Scanning tunnelling microscopy studies of vapour deposited films of tetrathiafulvalene with iodine, Appl. Phys. Lett., 62 1074± 1076. Brodie I and Muray J J (1992) The Physics of Micro/Nano-Fabrication, Plenum Press, New York. Bullen D, Chung S-W, Wang X, Zou J, Mirkin C A and Liu C (2004) Parallel dip-pen nanolithography with arrays of individually addressable cantilevers, Appl. Phys. Lett., 84, 789±791. Cheung J H, Stockton W B and Rubner M F (1997) Molecular-level processing of

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conjugated polymers. 3. Layer-by-layer manipulation of polyaniline via electrostatic interactions, Macromolecules, 30, 2712±2716. Chi L F, Eng L M, Graf K and Fuchs H (1992) Structure and stability of LangmuirBlodgett films investigated by scanning force microscopy, Langmuir, 8, 2255±2261. Chrisey D B and Hubler G K (eds) (1994) Pulsed Laser Deposition of Thin Films, WileyInterscience, New York. Critchley S M, Willis M R, Cook M J, McMurdo J and Maruyama Y (1992) Deposition of ordered phthalocyanine films by spin-coating, J. Mater. Chem., 2, 157±159. Davies G H and Yarwood J (1989) Infrared intensity enhancement for Langmuir-Blodgett monolayers using thick metal overlayers, Langmuir, 5, 229±232. Decher G and Schlenoff J B (eds) (2003) Multilayer Thin Films Sequential Assembly of Nanocomposite Materials, Wiley-VCH, Weinheim. Dimitrakopoulos C D, Furman B K, Graham T, Hegde S and Purushothaman S (1998) Field effect transistors comprising molecular beam deposited , !-di-hexylhexathienylene and polymeric insulator, Synthetic Metals, 92, 47±52. Eden J G (1992) Photochemical Vapour Deposition, Wiley-Interscience, New York. Evenson S A, Badyal J P S, Pearson C and Petty M C (1996), Variation in intermolecular spacing with dipping pressure for arachidic acid LB films, J. Phys. Chem., 100, 11672±11674. Gimzewski J (1998) Molecules, nanophysics and nanoelectronics, Physics World, 11, No. 6, 25±27. Grigore L and Petty M C (2003) Polyaniline films deposited by anodic polymerization: properties and applications to chemical sensing, J. Materials Science Materials in Electronics, 14, 389±392. Hara M and Sasabe H (1995) in Petty M C, Bryce M R and Bloor D (eds) An Introduction to Molecular Electronics, Edward Arnold, London. Hong S and Mirkin C A (2000) A nanoplotter with both parallel and serial writing capabilities, Science, 288, 1808±1811. Hong S, Zhu J and Mirkin C A (1999) Multiple ink nanolithography: toward a multiplepen nanoplotter, Science, 286, 523±525. Hong C, Chae H B, Lee K H, Ahn S K, Kim C K, Kim T W, Choi N I and Kim S O (2002) The possibility of pulsed laser deposited organic thin films for light-emitting diodes, Thin Solid Films, 409, 37±42. Kastantin M J, Li S, Gadre A P, Wu L Q, Bentley W E, Payne, G F, Rubloff G W and Ghodssi R (2003) Integrated fabrication of polymeric devices for biological applications, Sensors and Materials, 15, 295±311. Kinbara A, Hayashi T, Wakahara K, Kikuchi N, Kusano E and Nanto H (2003) Polyimide-based organic thin films prepared by rf magnetron sputtering, Thin Solid Films, 433, 274±276. Lee K B, Park S J, Mirkin C A, Smith J C and Mrksich, M (2002) Protein nanoarrays generated by dip-pen nanolithography, Science, 295, 1702±1705. Liu X, Fu L, Hong S, Dravid V P and Mirkin C A (2002) Arrays of magnetic nanoparticles patterned via `dip-pen' nanolithography, Adv. Mat., 14, 231±234. Lucovsky G, Ibbotson D E and Hess D W (eds) (1990) Characterization of PlasmaEnhanced CVD Processes, Materials Research Society, Pittsburg. Lvov Y, Onda M, Ariga K and Kunitake T (1998), Ultrathin films of charged polysaccharides assembled alternately with linear polyions, J. Biomater. Sci. Polymer Edn., 9, 345±355.

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Maggioni G, Carturan S, Quaranta A, Patelli A, Della Mea G and Rigato V (2003) Deposition of fluorescent organic coatings by glow discharge induced sublimation, Surface & Coatings Technology, 174, 1151±1158. Maissel L I and Glang R (eds) (1970) Handbook of Thin Film Technology, McGraw-Hill, New York. Maruccio G, Cingolani R and Rinaldi R (2004) Projecting the nanoworld: concepts, results and perspectives of molecular electronics, J. Mater. Chem., 14, 542±554. Netzer L and Sagiv J (1983) A new approach to construction of artificial monolayer assemblies, J. Amer. Chem. Soc., 105, 674±675. Netzer L, Iscovici, R and Sagiv, J (1983) Adsorbed monolayers versus LangmuirBlodgett monolayers ± why and how? I: from monolayer to multilayer, by adsorption, Thin Solid Films, 99, 235±241. Nicklow R M, Pomerantz M and Segmhller A (1981) Neutron diffraction from small numbers of Langmuir-Blodgett monolayers of manganese stearate, Phys. Rev. B, 23, 1081±1087. Park J H, Kang C H, Kim Y J, Lee Y S and Choi J S (2004) Characteristics of pentacenebased thin-film transistors, Mat. Sci. Eng. C, 24, 27±29. Paul K E, Wong W S, Ready S E and Street R A (2003) Additive jet printing of polymer thin-film transistors, Appl. Phys. Lett., 83, 2070±2072. Pearson C, Nagel J and Petty M C (2001) Metal ion sensing using ultrathin organic films prepared by the layer-by-layer adsorption technique, J. Phys. D: Appl. Phys., 34, 285±291. PercËin G and Khuri-Yakub B T (2002) Micromachined droplet ejector arrays for controlled ink-jet printing and deposition, Rev. Sci. Inst., 73, 2193±2196. Peterson I R and Russell G J (1984) An electron diffraction study of x-tricosenoic acid, Philos. Mag. A, 49, 463±473. Peterson I R, Russell G J, Earls J D and Girling I R (1988) Surface pressure dependence of molecular tilt in Langmuir-Blodgett films of 22-tricosenoic acid, Thin Solid Films, 161, 325±331. Petty M C (1996) An Introduction to Langmuir-Blodgett Films, Cambridge University Press, Cambridge. Petty M C (2000) in Functional Organic and Polymeric Materials, ed. T. H. Richardson, Wiley, Chichester. Petty M C, Bryce M R and Bloor D (eds) (1995) An Introduction to Molecular Electronics, Edward Arnold, London. Piner R D, Zhu J, Xu F, Hong S and Mirkin C A (1999) `Dip-pen' nanolithography, Science, 283, 661±663. Rees I D, Robinson K L, Holmes A B, Towns C R and O'Dell R (2002) Recent developments in light-emitting polymers, MRS Bulletin, 27, 451±455. Richardson T H (ed.) (2000) Functional Organic and Polymeric Materials, Wiley, Chichester. Salih A J, Marshall J M and Maud J M (1997) High-mobility low-threshold pentacene thin-film transistors prepared at rapid growth rates by pulsed-laser deposition, Phil. Mag. Letts., 75, 169±177. Sazaki G, Fujino T, Sadowski J T, Usami N, Ujihara T, Fujiwara K, Takahashi Y, Matsubara E, Sakurai T and Nakajima K (2004) Epitaxial relation and island growth of perylene-3,4,9,10-tetracarboxylic dianhydride (PTCDA) thin film crystals on a hydrogen-terminated Si(111) substrate, J. Crystal Growth, 262, 196±201.

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Sherman A (1987) Chemical Vapour Deposition for Microelectronics, Noyes Publications, Park Ridge. Sirringhaus H, Kawase T and Friend R H (2001) High-resolution ink-jet printing of allpolymer transistor circuits, MRS Bulletin, 26, 539±543. Sirringhaus H and Shimoda T (2003) Inkjet printing of functional materials, MRS Bulletin, 28, 802±803. Speakman S P, Rozenberg G G, Clay K J, Milne W I, Ille A, Gardner I A, Bresler E and Steinke J H G (2001) High performance organic semiconducting thin films: ink jet printed polythiophene [rr-P3HT], Organic Electronics, 2, 65±73. Stockton W B and Rubner M F (1997) Molecular-level processing of conjugated polymers. 4. Layer-by-layer manipulation of polyaniline via hydrogen-bonding interactions, Macromolecules, 30, 2717±2725. Thomas P J, Kulkarni G U and Rao C N R (2004) Dip-pen lithography using aqueous metal nanocrystal dispersions, J. Mater. Chem., 14, 625±628. Tour J M (2003) Molecular Electronics, World Scientific, New Jersey. Tredgold R H (1994) Order in Thin Organic Films, Cambridge University Press, Cambridge. Ulman A (1991) An Introduction to Organic Thin Films, Academic Press, San Diego. Ulman A (ed.) (1994) Characterization of Organic Thin Films, Butterworth-Heinemann, Boston. Xia C, Fan X W, Park M K and Advincula R C (2001) Ultrathin film deposition of polythiophene conjugated networks through a polymer precursor route, Langmuir, 17, 7893±7898. Yasuda H (1985) Plasma Polymerization, Academic Press, New York.

4

Membranes and permeable films

N A H O E N I C H , University of Newcastle upon Tyne, UK and D M A L I K , Loughborough University, UK

4.1

Introduction

Membranes and thin films represent a barrier between bulk phases which permit selective mass and fluid transport by one or more driving forces, e.g., concentration gradient, pressure difference or electrical potential. Membranes and thin films may be manufactured from a variety of materials and employed in a wide range of applications such as membrane bioreactors for the production of biopharmaceuticals including macromolecular products from fermentation and cell culture, protein recovery and concentration as well as use in clinical settings for treatment of disease or as part of analytical devices, e.g., biosensors. For a number of years, membranes were considered as inert phase separators, however, when used in a clinical context this approach is no longer appropriate as the membrane materials interact with biological fluids. The purpose of this chapter is to provide an overview of membranes and thin film materials and to discuss their manufacture, application and functional performance including biocompatibility.

4.2

Materials and applications

4.2.1 Materials Within the human body, the plasma membrane (Fig. 4.1) regulates the movement of substances into and out of the cell. In order for species to survive, a constant influx and efflux of materials in and out of the cell occurs. There are a number of mechanisms by which substances cross cell membranes: passive diffusion through the lipid matrix, diffusion through protein channels, and active transport. In membrane separation processes these functions are replicated but in a much less sophisticated form. It is potentially possible to manufacture membranes and thin films from a variety of precursor materials including synthetic organic polymers, natural polymer materials such as cellulose, ceramic materials or metals (Fig. 4.2). The manufacturing process for membranes and thin films is varied. The method used will be determined by

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4.1 The structure of plasma membrane. (a) shows the basic arrangement of the lipid bilayer, and (b) a simplified model of the membrane protein arrangement. From Human Physiology The Basis of Medicine by Gillian Pocock, and Christopher D Richards (1999) reproduced by permission of Oxford University Press.

4.2 View of a 0.2 m ANOPORETM inorganic membrane showing 0.23 m latex microspheres retained on the membrane's surface. Membrane thickness approximately 60 m. Illustration courtesy of Whatman Plc.

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the format of the membrane, (sheet, tubular or hollow fibre) as well as its application. For example, the manufacture of membranes based on cellulose is by a regeneration process involving a solution of purified cellulose in sodium hydroxide (Visking Process) or by the solubilisation of cellulose in an ammonia solution of cupric oxide (Cuprammonium process). These methods produce membranes that are essentially a macroscopically homogeneous material that is hydrophilic and forms a hydrogel when in contact with liquid. Modified cellulose membranes are a subset of membranes based on cellulose in which modifications aimed at the reduction of the number of hydroxyl (OH) groups in the polymer with other less reactive materials. Substitutions used include acetyl groups (to produce cellulose acetate) N,N, diethylaminoethyl (DEAE) the replacement of the benzyl groups by ether groups or by the grafting of polyethyleglycol (PEG) to the membrane to form a hydrogel on the material's surface. The most recent development of modified cellulose membranes has been the coating of the membrane with Vitamin E, (d-tocopetherol) aimed at reducing the oxidative stress produced during extracorporeal circulatory procedures. Membranes manufactured from synthetic polymers tend to be hydrophobic and require blending with hydrophilic agents generally polyvinylpyrrolidone (PVP). Membranes based on synthetic materials are more porous than their cellulose counterpart, and differ significantly in their structure from cellulosebased materials. The overall structure of the synthetic membranes is dependent upon the manufacturing process and the ratio of the blended polymers. Several different techniques of manufacture exist: sintering, in which the porous structure is derived from treatment of granular material at a high temperature and pressure, stretching, in which the structure of an extruded film made from a crystalline material is modified, etching, whereby the film is exposed to highenergy particles produced by a radiation source and phase inversion. In phase inversion, the polymer is transformed in a controlled manner from a liquid to a solid state. Phase inversion covers a range of different techniques including immersion precipitation and can be used to produce either sheet, tubular or hollow fibre membranes, the latter being produced by the use of a spinnerette as opposed to a flat casting knife. The polymer solution, containing the base polymer plus a solvent and/or other additives such as a second polymer, is pumped through a spinerette under pressure, the extruded fibre then passes through a series of baths where coagulation and structure formation occurs. Both cellulose based membranes and membranes based on synthetic polymers, such as polysulfone, polyarylethersulfones and polyethersulfones, and polyamide are used widely in clinical and industrial applications. In contrast to cellulose membranes which may be considered as a simple hydrogel, the molecular transport in synthetic membranes occurs within the thin surface layer (the skin), whilst the substructure acts as a mechanical support. It is important to bear in mind that whilst the same polymer base might be used, the

Membrane type and pore radius

Symmetric microporous, 100±10,000 nm

Asymmetric microporous, 1±10 nm

Asymmetric skin-type, 0.5±1.5 nm

Cation and anion exchange membrane

Asymmetric homogeneous polymer

Asymmetric homogeneous polymer (A non-porous membrane)

Thin-film membranes

Process

Microfiltration

Ultrafiltration

Reverse osmosis

Electrodialysis

Gas separation

Pervaporation

Nanofiltration

Cellulose acetate and aromatic polyamide

Polyacrylonitrile, polymers

Polymers and copolymers

Sulfonated cross-linked polystyrene

Polymers, cellulosic acetate, aromatic polyamide

Polysulfone, polypropylene, nylon, PTFE, PVC, acrylic copolymer

Cellulose nitrate or acetate, polyvinylidene difluoride (PVDF), polyamides, polysulfone, PTFE, metal oxides

Membrane material

930±1590 kPa

Vapour pressure gradient

Hydrostatic pressure and concentration gradients

Electrical potential gradient

Hydrostatic pressure difference at approx. 2±10 mPa

Hydrostatic pressure difference at approx. 0.1±1.0 mPa

Hydrostatic pressure difference at approx. 10±500 kPa

Process driving force

Removal of hardness and desalting

Separation of azeotropic mixtures

Separation of gas mixtures

Desalting of ionic solutions

Separation of salts and microsolutes from solutions

Separation of macromolecular solutions

Sterile filtration, clarification

Applications

Table 4.1 The characteristics of membranes used in different membrane separation processes, process driving forces and applications of such processes

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manufacturing process will influence the structural morphology of the finished product which in turn will influence the material's functional and blood contact behaviour causing membranes based on the same polymer from different manufacturers to behave in quite different ways.1

4.2.2 Applications Membranes and thin films are produced in three basic configurations: sheet, tube and hollow fibre and are additionally used as encapsulation materials, e.g., encapsulating microspheres for use in controlled release drug delivery. The scale of applications range from laboratory, to industrial (Table 4.1). In addition membranes are also used extensively in clinical applications (Table 4.2). In this area the most widely used application is for the treatment of acute and chronic renal failure by haemodialysis, a diffusion based system in which metabolites elevated as a consequence of renal insufficiency are removed across a semi-permeable membrane contained in an artificial kidney or haemodialyser. The diffusion takes place in an electrolyte solution flowing on the outer side of the membrane. Haemodialysis favours the removal of smaller molecular weight solutes and for a number of years the traditional approach to assessing performance has been in terms of urea removal. A number of complications in patients with renal failure treated by haemodialysis may be ascribed to inadequate removal of `middle' molecules.2 This has led to the development of alternative treatment modalities in which species of molecules are transported across the membrane by either convection alone or by a Table 4.2 Medical devices utilising synthetic polymer membranes (adapted from Scott, K. 1995, Handbook of industrial membranes, 1st edition. Elsevier Science Publishers, Ltd) Environment

Device

Permeable solute

Ex vivo

Oxygenator Haemodialyser Haemofilter Plasma separator Plasma fractionator

o

O2, CO2 Small molecules, ions, middle molecules, H2O Plasma Plasma proteins

In vivo

Contact lens Biosensor Bioreactor Drug delivery systems Wound dressing Artificial grafts

o

O2 Biologically active molecules Pharmaceuticals O2, H2O Physiological solution, cells

In vitro

Blood filter Plasma separator Substrate for cell culture

Blood except WBC Plasma Physiological solution

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mixture of convection and diffusion. Separation occurs on the basis of relative size of the permeating species.3 The ability of such techniques to remove protein bound substances remains limited and for removal of such compounds alternative approaches must be used.4 Membranes can also be used in plasma separation or plasmapheresis. In plasmapheresis mediators of the disorder are removed by the filtration of blood against a membrane which retains the cellular elements of blood but removes the plasma component. The removed plasma component is then either treated or replaced by a mixture of fresh frozen plasma and electrolyte solution, infused to replace that removed. A variant of this approach is plasma fractionation in which, rather than replacing the plasma, the patient's own plasma undergoes further filtration to retain blood coagulation factors and immunoglobulins. Such techniques may also be used for the removal of low density lipoproteins.5 Membrane based devices also provide support during cardiac surgery in the form of extracorporeal blood oxygenators where gas transfer takes place into the patient's blood across a permselective porous membrane.6 Other clinical applications of membranes and thin films include the treatment of diabetes by the use of an artificial pancreas, whereby islets of Langherans are sandwiched between membrane sheets or grown inside fibres, thereby providing a degree of immuno-protection and at the same time permitting the transport of key components into the bloodstream that flows on the outer side of the membranes. Such an approach is also used in the treatment of acute renal failure utilising tubular cells deposited on a membrane matrix.7,8 The laboratory based applications of membranes and thin films focus on the separation of biomolecules or the viral reduction of blood products. As well as systems suited for handling large volumes of liquid solutions, samples of up to 500 microlitres for micro-purification of proteins, antibodies and nucleic acids may be achieved by the use of small micro-concentrator modules which can be placed into a centrifuge and spun. The centrifugal force generates forces the filtrate across the membrane and concentrates the sample. Membrane separation in the laboratory may utilise a feed stream which is forced through the membrane and collected. An example of `dead end filtration' is the disposable syringe end filters that utilise membranes bonded to a sealed plastic support plate. The fluid to be filtered is contained in the syringe and is manually driven through the filter and collected. Stirred cells are also examples of `dead end filtration'. In such devices a flat sheet of membrane is placed on an appropriate support structure and spacer that forms part of the cell assembly. The fluid is added and agitated by a magnetic stirrer. The cell is pressurised either by air or nitrogen and the filtrate is forced through the membrane. Alternatively a cross flow configuration may be used to facilitate molecular removal from a solute.

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Cross-flow systems are inherently more complex, but their use is associated with high rates of mass transfer, due to high shear rate or tangential velocity as well as turbulence adjacent to the membrane, both of which minimise membrane fouling. In cross-flow filtration modules, which may be hollow fibre or parallel plate design, the feed flow is parallel to the membrane and perpendicular to the filtrate flow. In both these modalities the solvent and solutes smaller than the pore size of the membrane pass through the membrane to form the permeate. Cross-flow systems can deal with a molecular weight range of 500 to 1,000,000 daltons or approximately 10 to 100 Angstroms and is suitable for concentrating or washing macromolecules such as proteins, peptides, polysaccharides, lipids, synthetic polymers, viruses and colloids or the concentration of diluted samples of nucleic acid. Thin films are generally used for coating purposes, for example, to facilitate controlled drug delivery. In such an approach, a polymeric membrane encapsulates the drug which is carried in the form of a saturated aqueous solution containing undissolved particles of the drug to maintain saturation. A variant of this concept is the use of a matrix device, which is analogous to a sponge. The polymer matrix is initially loaded with a drug saturated solution, maintained by drug particles as before. When such a device is implanted, the drug molecules diffuse across the polymer tissue interface and a drug depleted layer forms on the matrix side of the interface. As diffusion continues, this layer increases whilst the surrounding tissue concentration remains constant due to the washing effect of the body fluids in contact with the matrix.

4.3

Membrane characterisation

4.3.1 Physical structure A number of different approaches to membrane classification based on structure exist. Techniques related to permeation parameters (such as liquid and gas flux measurements, solute retention test, and perm porometry) characterise the thin layer in asymmetric membranes but they fail to provide insight into the underlying membrane structure. Morphology-related techniques (such as gas adsorption-desorption, atomic force microscopy, electron microscopy, mercury porosimetry, thermoporometry, and X-ray scattering) give more complete information on the porous structure. Atomic force microscopy (AFM), allows the surface study of nonconducting materials with resolution in the micrometre to subnanometre range, and is able to be used without special sample preparation. In addition to imaging, the technique can also be applied to investigate membrane electrical and mechanical properties, and adhesiveness.

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4.3.2 Transport processes A membrane interposed between two bulk phases or mixtures control the exchange of mass between them. The exchange process can be driven by different forces: a concentration difference across the membrane (dialysis), a pressure difference over the membrane (ultrafiltration, reverse osmosis) or by an (externally) applied electrical field (electrodialysis). If the membrane is impervious to molecular species then fluid alone may be transferred. The separation characteristics of membranes are often expressed in terms of their nominal molecular weight cutoff (NMWC) determined experimentally by measurement of the trans membrane transport of molecules with different sizes or with a range of solutes. The NMWC value is defined as equal to the molecular weight at which >90% is rejected. In cell membranes, transport across the membrane occurs as a consequence of passive transport driven by the kinetic energy of the molecules being transported or by membrane transporters. Thus a molecule or ion that crosses the membrane by moving down a concentration or electrochemical gradient and without expenditure of metabolic energy is transported passively (diffusion). For charged molecules, the electrical potential across the membrane also becomes critically important. Together, gradients in concentration and electric potential across the cell membrane govern passive transport mechanisms. In cell membranes, active transport also takes place, this being the movement of molecular against an electrochemical gradient which requires energy expenditure. In membranes used for membrane separation processes, the volume flux produced may be expressed from consideration of the hydraulic resistance of the membrane used and the pressure drop over the membrane. In general the inverse of the resistance, hydraulic permeability being used. The permeability is defined by the pore size and structure and thickness of the membrane. Membranes can have different structures, they may consist of different, more or less cylindrical pores or may be composed of a packed bed of various small particles (which are of the order of nanometres). If the membrane consists of pores, they are often cylindrical and straight. The permeability of such pores is usually described by the Hagen-Poiseuille equation for flow through cylindrical tubes. If the membrane consists of a packed bed of particles, the specific permeability is derived from the Carman-Kozeny equation. Usually the pressure difference is applied externally to achieve filtration. During this process, solutes are separated from the solvent. These solutes accumulate near the membrane interface. The concentrations of these solutes are thus higher at the feed side than at the permeate side. This causes an osmotic pressure difference that reduces the effective pressure difference across the membrane. The simultaneous presence of both concentration and pressure driving forces across the membrane results in a mixed diffusive convective transport process.

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When both these forces are significant, both must be accounted for as well as their interactions. Irreversible thermodynamics provides a framework for the analysis of these complex interactions originally in the form of Kedem Katchalsky equations. More recently a modification of this approach was described by Kargol and Kargol9 who derived mechanistic equations based on the classical approach. These equations have the form: JV ˆ LP
P ÿ LP 
 and  P
P JS ˆ !D
 ÿ …1 ÿ †CL and represent both the volumetric (V) and solute (S) fluxes. These equations encompass four different processes; filtration, (LP
P) osmosis, (LP 
) solute  P
P) and where Lp;  and diffusion (!D
) and solute convection (…1 ÿ †CL ! are coefficients of filtration, reflection or rejection and permeation,  is the average concentration in membrane pores. respectively, and C If fluid alone is being transported across the membrane, then the first of the above equations will apply. The volume flux generated as a consequence of ultrafiltration can also be expressed in terms of concentrations by:   CW ÿ CP J ˆ K log n CB ÿ CP where CW , CP , and CB are the concentration of solute at the membrane interface, in the permeate stream and in the bulk respectively, and K is the mass transfer coefficient in the boundary layer. The degree of separation of a solute from the solution is expressed in a rejection coefficient, which is defined as: 1ÿ

CP CW

The rejection coefficient is dependent upon the molecular species, and its size relative to the pore of the membrane. Membranes are unable to discriminate between macromolecules that have sizes similar to one another, and for this reason membrane separation processes are unsuitable for the separation of soluble molecules whose molecular weight or effective sizes in solution (Stokes radii) differ by less than a factor of 6. When filtering low concentrations of macromolecular solutions by ultrafiltration, their osmotic pressures are low compared to the applied pressure and can be neglected. The osmotic pressure exerted by such a solution is generally in the form:
 ˆ AC ‡ A1 C 2 ‡ A2 C 3 where A, A1 and A2 are osmotic virial constants and C is the concentration of the bulk macromolecular solution. The coefficient A describes Van't Hoff's limiting

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law for the osmotic pressure which is applicable at very dilute concentrations, whilst A1 can be expressed in terms of the solute molecular weight. Ion transport is important in both biological and biotechnological applications. Ion transport across the membrane will be influenced by charge on the surface of the membrane (zeta or streaming potential) and the manner in which the ions are transported across the membrane. Transport across uncharged membranes may be characterised by consideration of the movement induced by the electrical field, diffusion and bulk convection. The movement resulting from each of these forces may not be in the same direction. The analysis of ion transport across charged membranes is more complex as the charge distribution on the membrane affects the concentration and potential distributions.

4.4

Blood material contact

Membranes and thin films used in therapeutic processes involve their repeated contact with biological fluids such as blood. For example, the treatment of renal failure by dialysis is generally undertaken three times weekly and each treatment session lasts 3±5 hours. Membranes have traditionally been considered as inert barriers, however, it is increasingly recognised that interactions occur between the material and blood components (Fig. 4.3). Thus, in therapeutic treatments

4.3 Biological pathway activation following contact with a material.

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this aspect of membrane performance becomes important. A primary requirement, therefore, of any such material is that the material should be biocompatible. The classical definition of biocompatibility is: `the ability of a material to perform with an appropriate response in a specific application'. This historic definition was based on the principle that a biomaterial has to perform and not simply exist, that it has to be associated with appropriate responses to ensure satisfactory performance. This definition recognises the fact that the response to a material will vary from one situation to another and that the appropriateness may vary. Furthermore, the definition allows a distinction to be made between biocompatibility and biological safety. The main difficulty with this definition is that the applications of materials in the clinical setting are varied, and there may be little commonality with the appropriateness of the responses. To account for this, the original definition has recently been revised such that the biocompatibility of a medical device that is repeatedly in contact with blood may be considered as `the ability of the device to carry out its intended function within flowing blood, with minimal interaction between device and blood that adversely affects device performance, and without inducing uncontrolled activation of cellular or plasma protein cascades'.

4.5

Biological events at the membrane and thin film blood interface

The magnitude of biological interactions is governed by the material's surface rather than its bulk characteristics, i.e., molecular interactions. Parameters of importance include hydrophobicity, which may be calculated using contact angle measurements, the presence of specific chemical groups (X-ray photoelectron spectroscopy and IR-spectroscopy), surface roughness (optical profiles and scanning probe microscopy) as well as the flow conditions at the time of the interaction.10 The underlying disease condition and drugs administered, may also play a role. A number of the responses that occur are in part a consequence of the blood recognising the material as non-self, and in part a consequence of the interaction between the material and cell surfaces. In addition cellular components of blood may also be involved, such as the activation of platelets with subsequent generation of thromboxane, the release of leucocyte proteinase enzymes and the generation of cytokines and reactive oxygen species (ROS).

4.5.1 Protein deposition or adsorption There is a common belief in biomaterials research that cellular interactions with natural and artificial surfaces are mediated through adsorbed proteins. The initial event that takes place on blood material contact is the deposition of proteins onto

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the surface.11 This deposition is rapid, occurring in a few milli-seconds and is governed by the interfacial properties of the proteins and the material surface. The importance of protein deposition stems from the fact that the deposited proteins exert a strong influence on subsequent cellular interactions with the surface. The process of protein adsorption comprises the following steps: 1. 2. 3. 4. 5.

transport towards the interface attachment at the interface eventual structural rearrangements in the adsorbed state detachment from the interface transport away from the interface.

The initial proteins deposited may be displaced by other proteins in the following sequence: albumin, immunoglobulins, fibrinogen, fibronectin, high molecular weight kininogen (HMWK) and Factor XII. This is termed as the Vroman effect. The adsorption of proteins may be considered to provide a degree of bioreactivity to the membrane or film surface. On the other hand, in the case of porous materials, such deposition inevitably modulates the diffusive characteristics of molecular species of interest. The adsorption of proteins mediates cellular adhesion to the adsorbed proteins. Primarily platelets are involved, and this sensitivity may be explained by the presence of receptors on the platelet surface (IIb/IIIa, Ib/IX) which facilitate such adhesion and are also involved in neutrophil activation. Many different techniques have been applied to study the molecular adsorption of proteins at interfaces. Several reviews have extensively discussed these techniques.12 The most commonly used techniques for protein adsorption kinetics studies include optical and spectroscopic techniques and electrical methods. Optical techniques include ellipsometry, variable angle reflectometry and surface plasmon resonance (SPR). Ellipsometry can be used with both transparent and non-transparent substrates whereas SPR must be used with a noble metal substrate.13 Spectroscopic techniques rely on the interaction of photons with the species present in the interfacial region to detect molecular events at the interface and include infra-red absorption (IR), Raman scattering, fluorescence emission circular dichroism (CD) and optical waveguide lightmode spectroscopy (OWLS). OWLS is a highly sensitive technique allowing real-time monitoring of protein-substrate interactions based on the measurement of the polarisability density (refractive index) in the vicinity of the waveguide surface. Since radioactive, fluorescent or other types of tagging are not required, measurements reflect the behaviour of unmodified proteins at the biomaterial interface. Other techniques include the use of radioactive isotope labelled proteins to quantitatively determine the extent of the protein adsorption or the spatial

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distribution on the surface of the material. Scanning probe microscopy techniques such as AFM provide dynamic imaging of real-time protein adsorption or static imaging of the adsorbed proteins. More interestingly, quantitative measurements of the strength of molecular interactions between the protein and the biomaterial are now possible.14 Protein-biomaterial adhesion force is measured with an AFM using a protein modified probe tip to analyse the actual interaction forces between blood proteins and the biomaterial surface.15 Predicting quantitatively the adsorption of a given protein at a well characterised surface under specified solution conditions remains a challenge.

4.5.2 Activation of the coagulation system Activation of the coagulation system occurs as soon as the blood leaves the vessel, and comes into contact with a non-physiological surface. Minimisation of coagulation may be by the use of drugs with antiocoagulant properties such as Coumarin, Warfarin, glycosaminoglycans, heparin and heparan sulfate. Plasminogen activators may also be used for the control of coagulation. When a foreign surface comes into contact with blood, the preotein deposition onto the surface includes Factor XII which becomes activated (Factor XIIa) and initiates the clotting or coagulation cascade. Fibrinogen is also present on the surface causing the attachment of platelets. The surface aggregated platelets encourage more fibrinogen to fibrin conversion by the production of additional thrombin and thus providing a positive feedback loop up regulating the coagulation system. The thrombotic potential of a material may be most readily assessed in a static experiment in which it is compared with a reference material, usually silicone coated soft glass. The thrombotic potential of the surface is generally evaluted by the timing of the development of a clot. These measurements may be combined with the determination of hemolysis.16 This type of approach, however, does not take into account flow across the surface which may influence the rate of clot formation or red cell damage. In-vivo testing of new materials poses problems, and often an indication of the behaviour of the material may be elicited by the consideration of ex-vivo tests. In such tests, a small volume of blood drawn directly from an animal of volunteer is passed over or placed into contact with the material and is then collected and analysed; considerable deviations from true in-vivo tests may be observed. When considering the behaviour of membranes used in extracorporeal procedures such as haemodialysis, additional issues are present. All therapeutic procedures utilise an anticoagulant, heparin, although less commonly low molecular weight heparin may be used. The adequacy of heparinisation during the procedure is generally measured by clotting time measurement. Of these, the measurement of activated clotting time is the

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simplest, however, it is only a broad non-specific indicator. A more accurate measure is the measurement of activated partial thromboplastin time (APTT). Heparin activity may be determined from the measurement of Factor Xa which is directly related to the amount of heparin present. Serial measurements provide an indicator of the loss of heparin due to adsorption or metabolism although in a clinical setting it is difficult to distinguish between these two. Commercially available kits are available for its determination. These measurements do not, however, provide any indication of the ability of the material to induce thrombus formation. A commonly used method for the measurement of this is by the measurement of thrombin antithrombin III complex.

4.5.3 Activation of the immune system The most thoroughly researched system activated following blood material contact has been the complement system. Activation of the human complement system following contact with a membrane or thin film is principally via the alternative pathway. It is initiated by the deposition of C3b on the material surface which with Factor B forms C3 and C5 convertases. These enzymes cleave the anaphylatoxins C3a and C5a from C3 and C5 by an autocatalytic process. Once in the circulation the C terminal arginine is removed and C3a des Arg and C5a des Arg are formed. These fractions are generally measured when membrane induced complement activation is studied. The cleavage of C5 by C5 convertase results not only in the production of C5a but also of C5b which initiates a macromolecular complex of proteins, the membrane attack complex, formed from C5b, C6, C7, C8, C9. (C5b-9). Complement products are an important mediator of the inflammatory response produced by membranes and are important markers for the biocompatibility of such materials. A wide range of clinical sequalae are associated with the generation of C3a, C5a and C5b-9 and virtually every blood cell type responds either directly via receptors, or indirectly via secondary mediators to these complement fractions. Implantation of devices in contact with blood may be associated with interactions between the implanted material and the host immune system. The aberrant state of monocyte and T-cell activation resulting from these host/device interactions is accompanied by two parallel processes, first, selective loss of Th1 cytokine producing CD4 T-cells through activation-induced cell death and secondly, unopposed activation of Th2 cytokine producing CD4 T-cells resulting in B-cell hyper-reactivity and dysregulated immunoglobulin synthesis through Th2 cytokines and heightened CD40 ligand-CD40 interactions. The net results of these events is that the patient develops progressive defects in cellular immunity and is at risk of infection, or allosensitisation.

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In characterising complement activation by biomaterials, older publications dealing with complement activation refer to a less sensitive measure, total complement (CH50), to quantify activation. As discussed above, the primary event is the interaction of the C3 molecule with the material surface. This is followed by a complex series of cascade reactions during which C3a and C5a are released. Upon release into the circulation there is removal of the C-terminus arginine from C3a (C3a des Arg). Such modification abolishes anaphylotoxin activity but does not affect the platelet-stimulating activity of the peptide or release of serotonin.17 The circulating C3a des Arg levels may be measured by commercially produced ELISA kits, but currently no measurement of C5a des Arg is possible. Commercially produced ELISA assays are also available for the quantification of the membrane attack complex (C5b-9).

4.5.4 Cellular activation Leucocytes White blood cells are made up of neutrophils, eosinophils, basophils monocytes and lymphocytes. The principal functions of neutrophils and monocytes are their ability to recognise foreign material and in the body's response to infection. They are attracted to sites of inflammation and phagocytose foreign particles. Following phagocytosis, neutrophils degranulate. Macrophages share the ability of neutrophils to phagocytose foreign materials, but can ingest larger particles and deal with a greater number of particles. In therapeutic procedures such as haemodialysis, plasma separation and membrane oxygenation, following initial contact of the blood and foreign surface, a profound transient neutropaenia occurs. The neutrophils are known to aggregate in the lung vasculature, and it has been shown that the magnitude of neutropaenia is related to the degree of complement activation, however, the molecular mechanisms governing this feature of blood material contact remain incompletely elucidated. It is likely that the activation of neutrophils produced by C5a induces changes in the expression of leucocyte endothelial adhesion molecules as neutrophils have receptors for C3b, C5a, and integrins such as MAC 1 (CD11b/CD18) responsible for adhesion, the modulation of which leads to cell adhesion to the endothelium, and pulmonary sequestration. Measurement of such molecules may be by the use of ELISA kits or by the use of fluorescent antibody cell sorting (FACS) analysis. Neutrophil activation is also associated with superoxide production which can be measured following material contact. Granulocytes (polymorphonuclear neutrophils) release proteinases such as elastase, lactoferrin and myeloperoxidase which may be measured by ELISA kits.

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Platelets Circulating platelets adhere to proteins deposited on the material surface. Changes in circulating platelet numbers may be elucidated by serial measurements. The deposition, however, is not only dependent upon the material but also on the flow conditions. The deposited platelets on the material surface release their granule content. Granules may be alpha or dense. Alpha granules contain platelet factor 4 (PF4) and B thromboglobulin (BTG) Both may be used as markers of activation, however, in a clinical setting other factors need to be considered, namely the clearance of PF4 from the circulation by endothelial binding. Heparin may also induce the release of PF4 from endothelial cells. Release is governed by the material surface as well as circulating factors such as thrombin and thromboxane A2. The dense granules contain ADP, ATP pyrophosphate and serotonin. Kits are available for the measurement of PF4 and BTG. The up-regulation of surface markers on platelets such as CD62 following material contact can be quantified by FACS. Platelet neutrophil interactions Platelet leucocyte co-aggregate formation has been implicated in the pathogenesis of thrombosis and inflammation, and observed during haemodialysis.18 Such interactions may be of relevance in both haemostasis and inflammatory processes. In haemodialysis, the increased formation of platelet leucocyte micro-co-aggregates is thought to be related to a primary platelet activating mechanism that involves P selectin (CD62P) a marker of activated platelets as well as CD15s (the sialyl-Lewis x molecule) a selectin ligand and it is possible that the CD62P/CD15s interaction seen during haemodialysis represents the first stage of leucocyte margination. Monocytes Monocytes are activated during blood membrane contact. C5a induces mRNA IL- and TNF transcription and primes the cells for the translation of cytokines following further stimulation. This further stimulation may be endotoxin entering the bloodstream from the dialysis fluid across the membrane, or by the direct contact of the monocytes with the membrane itself. Eosinophils Eosinophils have anti-histamine properties and congregate around sites of inflammation. Normal blood contains between 1±6% eosinophils, and elevated numbers are known to be present in people with allergic conditions such as hay fever. Eosinophilia is known to occur in haemodialysis patients; its presence has

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been associated with allergy to dialysers and exaggerated activation of complement during HD. Its etiology, however, remains unknown. Erythrocytes Red cells are remarkably robust, however, mechanical damage may occur during extracorporeal circulatory processes arising from the blood pump or through rupture of the cell membrane by coming into contact with granular material. Lysis of red cells results in the release of ADP which is a potent mediator of platelet aggregation.

4.6

Improvement of biocompatibility

Improvement of the biocompatibility profile of a membrane material is generally achieved by the alteration of its bulk characteristics by alteration of the manufacturing process or the use of polymer blends. Increased understanding of the molecular mechanisms involved is resulting in better understanding of the role of material related properties influencing the molecular mechanisms and forms the basis of the development of materials with minimal bioreactivity. Controlled chemical modification can lead to the development of predictable biocompatibility profiles. For example, hydrophilic domains on the material surface in polymer blends have a stimulatory part on the complement activation potential of the material, but play little part in determining the material's ability to activate platelets. On the other hand, hydrophobic domains show a reduced influence on the activation of the complement system, but stimulate platelet adhesion.19 Plasma modification of the membrane surface may be a useful way to improve the biocompatibility of polymer membranes by utilising high-energy electrons, ions, atoms, radicals and excited molecules produced by electric discharge. The plasma generates chemically reactive species from otherwise inert molecules (at relatively low temperatures) that may be deposited on the membrane surface (changing its chemical composition) thereby affecting its biocompatibility (Table 4.3).

4.7

Conclusion

Although membrane separation processes are widely used in many major industries a number of unresolved technical requirements remain. These include the availibity of pre-evaporation membranes for organic-organic separations, oxidation-resistant membranes for reverse osmosis and ultrafiltration foulingresistant membranes. Electrodialysis is an established membrane separation process for the desalination of brackish waters which has changed little in the last ten years. Membranes with better temperature stability and spacers with

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Table 4.3 Biomedical applications of plasma modified polymer surfaces (adapted from Scott, K. 1995, Handbook of industrial membranes, 1st edition. Elsevier Science Publishers Ltd) Gases (or monomers)

Polymers

NH3 (or N2 and H2)

Polypropylene (PP) Poly (vinylchloride) (PVC) Polytetrafluoroethylene (PTFE) Polycarbonate (PC) Polyurethane (PU) Polymethylmethacrylate (PMMA)

Heparin bonding for improved blood compatibility

Hexamethyldisiloxane (HMDS) C2H4 + N2 Alllene + N2 + H2O

Poly (ethylene terephthalate) (PET) Silastic (SR) Polysulfone (PS)

Improved blood compatibility

C2H4, allene, styrene, acrylonitrile, C2F4, C2H3F, C2F3Cl, C2H3Cl

Polystyrene (PSt) SR

Improved tissue compatibility

C2H4, C2F3Cl styrene

SR

Improved tissue compatibility

C2H2 + N2 + H2O

PMMA

Modified contact lens wettability by proteins

improved flow distributions can produce incremental improvements in brackish water desalination systems. A number of novel technologies are also being developed, including magnetic separations, electrically driven systems and liquid membrane systems. In membranes used for clinical applications an ongoing focus is improvement of biocompatibility. New manufacturing techniques involving a more defined pore distribution have also been introduced.20 Tissue engineering offers a number of exciting opportunities for the treatment of a variety of medical conditions. The membranes and thin films provide an ideal artifical support structure for the development of bio-hybrid devices. Such devices are available for both kidney and liver support.21,22 Micro-encapsulation offers the potential to encapsulate the islets of Langherhans for the treatment of diabetes. Such encapsulation may also be used for affinity dialysis for the removal of HIV and toxic viral proteins from blood.23

4.8

References

1. Hoenich NA, Woffindin C, Brennan A, Cox PJ, Matthews JN, Goldfinch M (1996) A comparison of three brands of polysulfone membranes. J Am Soc Nephrol.; 7(6): 871±6.

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2. Vanholder R, Smet RD, Glorieux G, Dhondt A (2003) Survival of hemodialysis patients and uremic toxin removal. Artif Organs.; 27(3): 218±23. 3. Colton CK (1987) Analysis of membrane processes for blood purification. Blood Purif.; 5(4): 202±51. 4. Sauer IM, Goetz M, Steffen I, Walter G, Kehr DC, Schwartlander R, Hwang YJ, Pascher A, Gerlach JC, Neuhaus P (2004) In vitro comparison of the molecular adsorbent recirculation system (MARS) and single-pass albumin dialysis (SPAD). Hepatology; 39(5): 1408±14. 5. Klingel R, Fassbender T, Fassbender C, Gohlen B (2003) From membrane differential filtration to lipid filtration: technological progress in low-density lipoprotein apheresis. Therap Apher Dial.; 7(3): 350±8. 6. Wickramasinghe SR, Goerke AR, Garcia JD, Han B (2003) Designing blood oxygenators. Ann N Y Acad Sci.; 984: 502±14. 7. Isayeva IS, Kasibhatla BT, Rosenthal KS, Kennedy JP (2003) Characterization and performance of membranes designed for macroencapsulation/implantation of pancreatic islet cells. Biomaterials.; 24(20): 3483±91. 8. Humes HD, Weitzel WF, Fissell WH (2004) Renal cell therapy in the treatment of patients with acute and chronic renal failure. Blood Purif.; 22(1): 60±72. 9. Kargol M, Kargol A (2003) Mechanistic equations for membrane substance transport and their identity with Kedem-Katchalsky equations. Biophys Chem.; 103(2): 117±27. 10. Spijker HT, Graaff R, Boonstra PW, Busscher HJ, van Oeveren W (2003) On the influence of flow conditions and wettability on blood material interactions. Biomaterials.; 24(26): 4717±27. 11. Ramsden JJ (1995) Puzzles and paradoxes in protein adsorption. Chemical Society Reviews; 24: 73±78. 12. Ramsden JJ (1994) Experimental methods for investigating protein adsorption kinetics at surfaces. Q Rev Biophys.; 27(1): 41±105. 13. Voros J, Ramsden JJ, Csucs G, Szendro I, De Paul SM, Textor M, Spencer ND (2002) Optical grating coupler biosensors, Biomaterials; 23: 3699±3710. 14. Kidoaki S, Matsuda T (2002) Mechanistic aspects of protein/material interactions probed by atomic force microscopy. Colloids and surfaces B.; 23: 153±163. 15. Kidoaki S, Matsuda T (1999) Adhesion forces of the blood plasma proteins on selfassembled monolayer surfaces of alkanethiolates with different functional groups measured by an atomic force microscope. Langmuir; 15: 7639±7646. 16. American Society for Testing Materials (1997) ASTM F 756-93. 17. Polley MJ, Nachman RL (1983) Human platelet activation by C3a and C3a des-arg. J Exp Med.; 158(2): 603±15. 18. Gawaz MP, Mujais SK, Schmidt B, Blumenstein M, Gurland HJ (1999) Plateletleukocyte aggregates during hemodialysis: effect of membrane type. Artif Organs.; 23(1): 29±36. 19. Deppisch R, Gohl H, Smeby L (1998) Microdomain structure of polymeric surfaces ± potential for improving blood treatment procedures. Nephrol Dial Transplant.; 13(6): 1354±9. 20. Ronco C, Bowry S (2001) Nanoscale modulation of the pore dimensions, size distribution and structure of a new polysulfone-based high-flux dialysis membrane. Int J Artif Organs.; 24(10): 726±35. 21. Humes HD, Weitzel WF, Fissell WH (2004) Renal cell therapy in the treatment of

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patients with acute and chronic renal failure. Blood Purif.; 22(1): 60±72. 22. Patzer II JF, Lopez RC, Zhu Y, Wang ZF, Mazariegos GV, Fung JJ (2002) Bioartificial liver assist devices in support of patients with liver failure. Hepatobiliary Pancreat Dis Int.; 1(1): 18±25. 23. Tullis RH, Duffin RP, Zech M, Ambrus JL (2003) Affinity hemodialysis for antiviral therapy. II. Removal of HIV-1 viral proteins from cell culture supernatants and whole blood. Blood Purif.; 21(1): 58±63.

5

Stable use of biosensors at the sample interface J F G A R G I U L I , University of London, UK, A G I L L and G L I L L I E , University of Manchester, UK, M S C H O E N L E B E R , University of London, UK, J P E A R S O N , University of Manchester, UK, G K Y R I A K O U and P V A D G A M A , University of London, UK

5.1

Introduction

Biosensors provide a high degree of elegance in regard to their simple juxtaposition of a bioreagent and a transducer function. To work properly, there must be a direct alignment of a functionally responsive biolayer and a transducer element, which is able to directly extract the binding information resulting from the encounter with the analyte. There is a difficulty associated with such a simple, structurally inflexible combination, due to the fact that optimisation is limited as compared with say the use of a liquid phase bioreagent with its attendant optimised solution parameters of pH, pI, concentration and reagent additives. However, the net result is a solid-state monolithic structure with the potential to perform analyses and to operate potentially in optically opaque samples. In biomedicine, the latter capability holds considerable advantages. Indeed, most if not all biofluids contain colloidal materials that are liable to render the sample optically opaque or, at the very least, to induce a certain amount of light scattering. Furthermore, the majority of clinical sample assays rely on absorbance techniques. Therefore, it is necessary to have access to biosensors that would perform reliably even in opaque samples. Nonetheless, there are still a few drawbacks associated with the use of biosensors in vivo. Although the established biosensor systems operate on the `macro' scale and have seen varying degrees of clinical exploitation,1 a key reason for their limited introduction into the application domain has been the rapid alteration of the biosensor interface through the surface activity of the colloidal elements of any unmodified biological sample. This problem is also intricately linked to that of biocompatibility of the exposed surface of the biosensor in direct contact with either living tissues or the physiological fluids present in the body.

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5.2

Biosensor limitations

5.2.1 Biosensors and bioresponse compromise Whilst the bulk amount of protein, colloid and cell transfer to a `clear' biosensor surface may be relatively low in amount, its diffusional barrier effect on the continued flux of analyte to a responsive biolayer surface will immediately be registered as a reduced biosensor response. The function of any flux dependent biosensor, unless it is merely a qualitative registering device, will be affected sufficiently that accuracy and precision of measurement will be lost, and neither may be recovered simply by recalibration. The devices therefore likely to be most affected will be those using degrading enzymes rather than where a true binding equilibrium is approached, e.g., antibody, lectins, receptors and DNA/RNA. In the case of the latter group, the only effect should be on the rate of approach to equilibrium not on its final value, unless of course the fouling biolayer affects solute partitioning. This is possible in principle if, say, the fouling layer is a charged colloid, and the analyte target is also charged; electrostatic forces are then liable to come into play to either partition in or partition out a given analyte. Exclusion, furthermore, is possible if the analyte target is a macromolecule, in which case it may have limited access to the affinity surface of the biosensor through the limited permeability of the colloid fouling layer. The underlying transducer of the biosensor, whilst not directly affected by surface colloidal deposits, may register the presence of these non-specific elements through its detection domain. Thus, the evanescent wave of an optical wave-guide or SPR system will respond equally to non-specific binding as to specific affinity interactions. Passivation of the transducer element is possible with some surface-active crystalloids. Thus free amino acids and thiol-containing molecules that are surface active can distort and depress the catalytic behaviour of a Pt working electrode used for redox dependent biosensors.

5.2.2 Biosensors and the selectivity compromise Whilst the biological component of a biosensor has accepted selective properties, and is de facto the driver for biosensor development, the underlying transducer, whether based on electrochemical, potentiometric, optical or microgravimetric principles, is vulnerable to a false positive response due to the surface activity of analogue species of either the target molecule or a molecule that is part of the transduction cascade. The most potent expression of the problem is where an electrically polarised noble metal electrode is used to detect the H2O2 product of an oxidase enzyme catalysed reaction in an enzyme electrode. At the typical polarising voltage of +0.65 V vs. Ag/AgCl, numerous species in biological solution are simultaneously oxidased, and false positive responses therefore result.2

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5.2.3 Interfacial problems at microfabricated biosensors The issues of direct biosample interfacing of biosensors applies to all biosensors irrespective of length scale. So whilst the x-y plane architecture may be geometrically precise, formulated in a much more reproductive manner using MEMs and other microforming technologies, the biological response of the host sample reacts in rather similar ways, with adverse buildup along the z axis, i.e., normal to the sensor surface. One rider to this equivalence of macro/microsensor outcomes is where response is flux (continuous diffusion) dependent; a sufficiently small sensing microsurface will have spherical and not a planar diffusion based supply of analyte and is thereby less affected by external variables such as fouling. There is also evidence that a microstructure may set up a lower intensity tissue reaction thus leading to reduced surface fouling.

5.3

Biocompatibility

The term `biocompatibility' covers the whole range of interactions that exist between a biomaterial implant and its biological surroundings, as well as the orchestrated sequence of responses the body invokes to essentially reject that implant as non-self. In the case of invasive in vivo monitoring, an electrochemical biosensor requires intimate, direct contact with the sample matrix in order to function properly, notwithstanding the intensity of the body's reactive response to its constituent materials. As a consequence, the observed performance of the biosensor is highly vulnerable to the local accumulation of surface-active agents from the body such as cells, proteins and other less well identified constituents such as colloidal and lipid aggregates. This accumulation of biological compounds on the surface of the sensing device is known as `biofouling', and will with time alter and degrade the biosensor response and performance. Furthermore, biosensors are not bio-inert, not only because of the active redox components they may incorporate but also because of the polymeric materials, as well as the coated or uncoated metal and carbon electrode interfaces they present to the living tissues. In these tissues, they provoke a high intensity local inflammatory reaction simply because they inhabit a wound site and, in blood, they are a nidus for surface coagulation with the attendant threat of local microand later macro-thrombi. The danger of thrombus dissemination in the vasculature, thromboembolism, constitutes a particular concern because of the possibility of considerably wider distribution of tissue damage. There are many examples of implants leading to thrombosis, embolism and death. While the risk might be acceptable for a life-saving device such as a vascular stent, such risk cannot be tolerated when a diagnostic device such as a biosensor is being used. Even if an implanted material is reputed to have no interaction with a biomatrix and is therefore believed to be bio-inert, some interactions at a microscopic scale still occur. A completely biocompatible material does not

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exist. As a consequence, all available materials provoke, to some extent, a biological response, adverse or not, that will impact on overall sensor performance. Indeed, even air entrapped within a tissue will ultimately provoke a local body reaction. Additional bio-effects, beyond local surface phenomena, that need to be taken into account when designing a biosensor for clinical use are carcinogenicity, mechanical stability, immunogenicity, chemical stability and biomechanical compatibility with the local soft tissue. Also, sensors may need to operate in vivo for quite different periods of time, and this must be taken into account when assessing the tolerance limit for biocompatibility. Admittedly, in contrast to the carefully assessed quantitative analytical behaviour of sensors in vitro, little progress has been made in regard to in vivo performance standards and acceptability. Also, the actual limits of tolerance for the degradation of function, resulting from the body's response, are still poorly understood and known. In vivo biocompatibility can be regarded as a hybrid between biomaterials and biosensor research, which deals with both the specific, as well as the complex interactive and cumulative effects of sample matrix constituents upon sensor function and operational life-time.3±5 For all sensors, understanding and predicting the in vivo biological response still poses a great, unmet, challenge. The crucial point is that, in contrast to conventional biomaterials, the surface deposition of the body's diffusive and cellular biocomponents leads to degradation of function, observed within minutes and hours rather than days and months. Whilst the device continues to function, its value as an accurate and precise quantitative system is lost. A hierarchy of biological interfacial phenomena exist, which, though subclassified relatively easily, are difficult to unravel in relation to their complex dynamics and as concerted interactive phenomena. For a long time researchers attempted to avoid the bioresponse entirely via the `mythical' concept of the totally bio-inert implant, leading later to the emergence of the idea of functional biocompatibility.6 In this latter case, a primary issue is not so much the total avoidance of the bioresponse, but rather achieving sufficient control over its adverse effects on the biosensor performance. The definition of biocompatibility is itself an approximation of multiple concepts. However, it is based on the expectation that certain types of materials will be able to provoke just a limited body response. No matter what the material deposited and accumulated at a sensor surface may be, it is whether or not a degradation of response occurs that really matters. The counterpart to this is whether an implant material poses a threat to the patient or not. Both immunological and toxicological dangers exist, for example, through the leaching of additives within covering polymers, the release of polymer degradation products, of metal/metal oxide particulates or of carbon electrode constituents. All are capable of triggering an early inflammatory and toxicological response, but the dangers of long-term effects, including

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carcinogenicity are largely uncharted. Again such dangers may be acceptable in the instance of acute life-saving or life quality enhancing implants such as heart valves, aortic grafts and pacemakers, but in the case of a measurement modality such as glucose monitoring, any such dangers have to be negligible. A biosensor can be regarded as a composite material, usually made up of metallic, polymeric and biological components. While it constitutes a minor burden as an implant, as it exhibits trivial mass and volume in relation to the body, it is however more likely to present a multiple combination of potential leachables, which, together with the reaction products of the biosensing reaction, could have significant local effects. Underlining all this is the fact that any release of antigenic protein poses special dangers as an immunogenic trigger, especially if a biosensor is to be implanted repeatedly. The biocomponent (e.g. enzyme) may itself be associated with an undefined constituent due to the presence of low level impurities, as in any bioagent, rendering traceability difficult. A practical expression of this comes from the routine use of bovine serum albumin (BSA) as a crosslinking matrix for glucose oxidase in the early days of research.7 Due to actual and theoretical dangers of the bovine spongiform encephalitis (BSE) agent, such a protein source for enzyme immobilisation is no longer acceptable. In view of the heterogeneity of both sensor internal structure and of the multiple types of surfaces presented to the biomatrix, the body's response maybe quite different over the active sensing regions of the device vs. the support regions. This makes it difficult to determine a precise structure ± biocompatibility linkage.8 Materials considerations are also relevant to the duration of operation envisaged. Short-term monitoring in, say, the critically ill patient, demands quite a different level of materials requirement to the more robust, mechanically resistant devices that, moreover, would have to be well tolerated by the patient over the long term, especially during ambulatory monitoring. Polymer encapsulant degradation and metal corrosion set a limit to the long-term implantation of the sensor in the patient. Furthermore, not only are the body's responses cumulative but a rigid device, mechanically incompatible with local soft tissue, will not be well tolerated, due to the stimulation of local pain sensors. Over time these released constituents, due to the immune response, will be upregulated; so will be the effects of unavoidable products of enzyme reactions and redox centre/mediator components. Wherever a foreign body is lodged in the tissue or the blood compartment, it can also serve as a focus for infection and poses a problem over the long term. Microbial films may form on the surface and these have a powerful way of resisting antibiotic assault. Such films are known as significant contributors to hospital infection rates.9 The situation is compounded if a device is only partially inserted as with, say, a percutaneous glucose needle sensor. These are typically implanted in vascular, partly fatty tissue to only a 10±15 mm depth. Whilst less invasive, they provide a contact pathway for externally derived

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microorganisms. Important in this regard is the fact that the skin surface has its own microbiological flora and cannot be fully sterilised. Diffusible, low molecular weight solutes in biological fluids provide a nominally stable solution environment in vivo, given that this internal environment is designed for physiological stability. Certainly, near-neutral pH conditions normally prevail, and the concentration variation of background electrolytes is relatively small, at least with regard to the major ions such as sodium, potassium and calcium. As an indicator of the total solutes, osmolality in blood ranges quite narrowly between 280±295 mOsm/kg. With the inherent stability of the enzyme within an electrochemical biosensor and its reduced dependence upon pH and background ionic changes due to the immobilisation, one would expect minimal background solute effects upon bioelectrochemical sensor function. However, even small variations will cause some analytical imprecision, especially as sample dilution or other specimen manipulation are precluded in contrast to bench top analysis. The above considerations have been a basis for the use of, say, sampling tissue ultra-filtrate for glucose monitoring via mechanical suction through permeabolised skin.10 However, low molecular weight solutes may still influence working electrode response through adsorption, passivation and then direct surface activity. Such effects may be difficult to identify when conducting measurements in a complex medium such as blood or tissue combining macroand microsolute influences. The former are only external surface active; the latter are able to permeate the entire device. The adsorbed solute can induce degradation of metallic components, as described for electrodes used in electrical stimulation, where increased ionic release and accelerated corrosion have been reported11 with a further facilitating influence due to surface active proteins.12 The electrochemical reaction can also contribute to electrode dissolution, which is of relevance to long-term implantation.13 Beyond standard cylindrical geometries and wire-type devices, emphasis is shifting to MEMs-based devices. However, some implantable sensors based on MEMs constructs are vulnerable to hydration and water ingress. As a consequence, rigorous packaging and encapsulation are needed for long-term function as well as over the short term. Typically, in order to avoid leakage currents and extraneous responses, inorganic or polymeric packaging materials are required.14,15

5.3.1 Protein constituents In terms of the total colloid biomass, proteins comprise the most important surface-active constituents of biofluids other than viable (and dead) cells. Total concentration in plasma is around 80 g/l. Proteins have a special ability to adsorb physically at solid/liquid interfaces and through consequent denaturation to form relatively adherent, immobilised phases, showing only partial remodelling or

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recycling between the bulk sample and surface. The adsorbed proteins thus undergo both conformational and orientational changes at the surface. As well as progressively increasing in depth with time, through deposition of solution originated proteins, the entire protein `gel' layer grows unpredictably, modulated by both environmental perturbation (convective, shear, pressure) and inherent structural remodelling, especially evident near the growth surface.16 Again, the key issue is that in contrast to conventional biomaterials, surface deposition effects on a new electrode surface are rapid and accentuated. Whilst the device continues to function, its value as an accurate and precise quantitative system is almost immediately lost, a particular concern in the case of glucose monitoring given the strict control demanded in diabetes management. Protein deposition occurs within seconds of contact,17 but inherent film composition and individual component abundance are conditioned by material surface energy (hydropholicity/hydrophilicity), charge, charge density, surface profile and roughness, degree of molecular ordering, pendent group flexibility and crystallinity. In the case of a polymeric interface, the presence of trace components, plasticiser content and minor degrees of surface oxidation all drive particular types of protein adsorption. Albumin, in particular, is a dominant constituent at the interface, and outcompetes other proteins in plasma such as immunoglobulins, fibrinogen and kininogen. Its presence at a surface, furthermore, helps to reduce the deposition of other surface fouling components including cells, and it is regarded as a passivating protein. The dynamics of the adsorption, however, appears to be different at different surfaces. For example, on a hydrophobic surface, adsorption is a single-step process, whereas on a hydrophilic surface, it involves two steps. The resultant in either case, though, may be the formulation of an equivalent layer with similar passivating behaviour.18 This new view of protein deposition reduces the importance of the primary surface in driving biocompatibility outcomes. Ongoing competition between proteins for the surface leads to the remodelling of the deposited surface layer and, in the context of, say, fibrinogen adsorption from plasma, its transient surface prevalence and subsequent partial displacement (the Vroman effect) represents the outcome of the finite number of available sites on the surface.19 In previous studies, surfaces have been alkyl modified in order to promote albumin deposition, and therefore to reduce platelet adhesion.20 More recently, immobilisation of highly hydrophilic polymers such as poly(ethylene oxide) (PEO) has proved especially valuable in resisting protein deposition.21 The equivalent has also been achieved, for example, through glow discharge deposition using tetraethylene glycol dimethyl ether.22 More classically, heparin immobilised at surfaces has proved effective in resisting a special type of surface deposition, the binding and activation of coagulation components. In this respect, surface bound heparin has been used to good effect at an intravascular oxygen catheter sensor.23

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Heparin works through its anionic sulphonate and aminosulphate groups,24 so attempts have been made to sulphonate artificial polymers, e.g., polystyrene and polyurethane.25 Also, the combined effect of PEO with end-attached sulphates to give heparin-like properties, and the added benefits of PEO flexibility has led to an improved thrombo-resistant polyurethane.26 Polyurethane is already frequently used as an outer membrane barrier for glucose sensors, so this may have a direct application when used in a bioelectrochemical sensor. The fundamental drawback of an enzyme-based biosensor is that it requires continuous substrate flux to the enzyme layer for an ongoing response. This flux has to be purely substrate concentration (gradient) dependent, unperturbed by any newly formed diffusion barrier in, on or around the biosensor itself. The first of the membrane-packaged devices developed to control such adventitious biolayer diffusion limitation was the Clark pO2 polarographic electrode where an external gas permeable membrane served to protect the working electrode from protein and colloid deposition.27 For surface anticoagulation, a primary need is for the prevention of the binding of factor XII, as this is a key trigger in the initiation of the coagulation cascade and eventual deposition of the crosslinked fibrin mat at the surface. An intrinsic pathway and an extrinsic coagulation pathway, the latter being induced by tissue-derived factors, may combine to create an accelerated cascade of fibrin deposition (Fig. 5.1). However, a parallel system of surface-active proteins has been rather neglected. This is the complement system, which undergoes an ordered, sequential response to an artificial surface, eventually to trigger the production and the release of flammatory mediators from white cells.27 This complement activation takes place through one of two pathways, the classical and the alternative. The classical complement pathway is initiated by antigen-antibody complexes, by crystals or bacterial and virus surfaces if antibodies are absent, or by complexes between positively and negatively charged molecules such as those between heparin and protamine. The alternative pathway is not triggered by immune/antibody complexes, but can be initiated by any foreign material introduced into the body, including a biomaterial, lipopolysaccaride, polysaccharide, or bioorganism. It is activated by surfaces with particular chemical characteristics allowing fragment C3b of a larger protein C3 at the surface to initiate the assembly of an amplification system, C3 convertase, at the surface28 for further C3b deposition (Fig. 5.2). With regard to specific sequences, C3 convertase cleaves C3 to generate C3a, and a further fragment C3b. In the nascent state, the latter binds to the surface and augments the convertase enzyme further to amplify C3b deposition on the surface. A further reaction leads to the cleavage of available C5, with the production of a C5a fragment and also recruitment of a C5±C9 sequence of effectors and associated inflammatory changes. It is clear that substantial complement activation can lead to major organ dysfunction, though admittedly

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5.1 Surface coagulation cascade initiated either by surface contact (the intrinsic pathway) or by tissue factor (TF, the extrinsic pathway). The two pathways are eventually converting, forming a fibrin clot due to activation of thrombin on fibrinogen. Factor XIII will eventually convert fibrin clot into insoluble fibrin gel.

only following large scale blood/surface interactions as in haemodialysers.29 However, this complex cascade also requires to be considered as a possible contributor to local events at the biosensor surface. Surface amines and hydroxyls, in particular, react with C3 to form complexes, though there is uncertainty whether these are covalent bonds or whether electrostatic and hydrophobic interactions are important.30 Outcomes with regard to systemic effects also need to be unravelled.31 Local effects of relevance to a bioelectrochemical sensor might be abated through control of the surface presented, perhaps through surface heparinisation.32

5.3.2 Blood interfacing Blood consists of about 55% plasma by volume with about 45% solid particles including proteins and cellular elements, especially red blood cells, but with a

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5.2 Surface induced complement activation leading to accelerated complement protein deposition at a surface expressing amino/hydroxyl groups. Side cascade generation of C5±C9 is also indicated.

small white blood cell and platelet volume contribution. Red blood cells contain mainly haemoglobin, carry oxygen to the peripheral tissues and normally do not leave the circulation. White blood cells, however, are able to leave the vascular system to move towards any active disease tissue focus from, say, that due to microorganisms intrusion through a foreign body. The induced tissue disruption is `sensed' very acutely, no matter how localised. Additionally, platelets operate as a natural blood containment system, typically by forming a mechanical microplug at a vascular injury site, eventually producing a defined clot to contain blood in the circulation. The cellular elements of blood play a complex, cooperative role in the maintenance of the surface coagulation process initiated by soluble coagulation factors (vide supra). However, the deposited protein layer from blood at the sensing surface changes the material interface considerably, and it is this layer over which the cellular elements then begin to accumulate. Initially, the most important of these are platelets, which normally circulate in the blood in an inactive state, but are also the most labile of all the formed elements and therefore the most difficult to evaluate in physiological studies. Once they come into contact with a foreign surface, they show strong adhesion. This spreading and aggregation behaviour, associated with the release of intra-platelet adenosine diphosphate (ADP), promotes secondary platelet aggregation and the creation of (adherent) thrombus. In the final analysis, the level of initial coagulation protein deposits at a surface are seen to correlate with platelet

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surface activity and therefore, both condition the blood compatibility of a given material.33 Quite apart from any surface effects, platelets are exquisitely sensitive to shear force. Both the magnitude of shear and the overall blood flow profile near a surface, through influence on platelet transport, can completely override surface affinity interactions. Moreover, turbulent conditions and their associated high shear are particularly able to trigger platelet activity and augment surface delivery.34 The platelet delivery process is further accentuated by red cell interactions and local fluid entrainment around red cells, promoting surface collisions.35 The net negative charge on a platelet, due to surface sialic acid groups, is important to surface attachment. However, surface receptor interactions are also a powerful determinant, such as the receptor mediated attachment of platelets to fibrinogen. Following adhesion, platelets degranulate with the release of ADP. Then, a host of other bioactive components are able to promote further platelet activation and the eventual creation of an adherent micro-thrombus. Incorporation into the thrombus of the other cellular components of blood, including white cells and red cells (both influenced by blood flow and pressure gradients), leads to thrombus growth. As a consequence, local blood flow is distorted near the originally smooth surface, further stimulating the growth of thrombus and aggregation of blood components of all types. The effect on a sensing surface is the reduction of solute transport to that surface and the local consumption of oxygen/glucose, depending upon the metabolic state of the cellular aggregates, thereby inducing changes on the locally measured parameters. It becomes especially difficult to characterise material-induced artefactual influences upon a sensor, let alone to model them. Therefore, their subtraction from true responses in vivo becomes problematic, demanding at the very least a frequent in-vivo calibration regimen. It may be that, in the future, the quantity and nature of a surface coagulation phase can be measured using an independent technique. Impedance measurements promise in this regard to allow at least for baseline drift36 but, in reality, the end result at present is the highly unsatisfactory need to recalibrate frequently. The imposed structure of an intravascular electrode disrupts blood laminar flow patterns and can be an indirect cause of surface thrombus formation. Experience with thrombolic events and clinical use of intravascular catheters indicates a relationship between vessel cross-section diameter and catheter size.37 Complicating the assessment of protein deposition, coagulation/ complement activation and platelet retention,38,39 surface irregularity and surface microdepressions over a device may promote thrombus formation.40,41 A smooth surface is an advantage, but, probably, surface anticoagulation is a more effective basis for reducing thrombus formation. There is a clear lesson here for the combination design of smooth, haemodynamically acceptable glucose sensor surfaces and their surface modification.

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Heparin is a natural anticoagulant in blood, and is the most frequently used bioactive surface ingredient used to reduce clotting. The protective action of heparin is based upon its stimulatory effect on antithrombin via an heparinantithrombin complex, though this may be countered by the fibrin interactions of thrombin.42 Surface heparinisation of membranes has allowed more reliable oxygen monitoring.43 Catheter heparinisation has also proved to be of general effectiveness.44 Covalent binding of heparin, though permanent, leads to a more rigid attachment which can reduce effectiveness so bridging groups are an advantage. Also, depending upon the required duration of sensor operation, heparin leaching from a porous membrane or other reservoir could lead to a more potent anticoagulation surface. Overall, heparin can play a major role in reducing the effects of coagulation, but it is not the complete solution to the problem hoped by some. Alternatives to heparin include low molecular weight anionic analogues and, in particular, poly(ethylene oxide) (PEO).45 Furthermore, the use of copolymer structures rather than surface attachment may provide a family of new blood-stable materials in the future.46

5.3.3 Tissue interfacing Subcutaneous tissue is a safer alternative to the intravascular siting of glucose sensors, avoiding the dangers of thromboembolism, as well as of rapid dissemination of infection. Problems of reliable monitoring, though different from those of blood, are nevertheless substantial. The implanted device, through its intrusion into a normal tissue architecture, is perceived by the body for what it is, a disruptive foreign body. As a response, the tissue sets up an intense (acute) inflammatory response designed to degrade, isolate and ultimately reject the foreign material.47 The outcome is locally distorted body fluid composition, i.e., modified functional physiology. As a consequence, no matter how reliable or biocompatible the sensor may be, measurements are thereby performed in an environment that is metabolically distorted and also appears to lose the rapid equilibrium relationship with the local blood and the capillary bed supply (Fig. 5.3). With a low-reactivity material, the acute inflammatory stage subsides to give way to tissue repair, which tends to generate a more vascular environment through capillary proliferation, as well as a matrix rich in fibroblasts. Capillary density in the vicinity of implanted electrodes has been shown to increase with a vascular density maximum at a 50 m distance from the sensor.48 However, measurement distortion was observed (sensor instability and slow response) and was thought to be due to local fluid increase. Whatever the mechanism, the distorted response was only marginally diminished for a tissue-implanted glucose sensor by using slow release of dexamethasone as an anti-inflammatory.49 Due to the presence of soluble bioactive agents, the tissue environment becomes highly hostile to the sensor. Local increases in hydrolase enzyme

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5.3 Three general phases of tissue response at implanted electrodes showing various local reactive phases.

activity (acid phosphatase, alkaline phosphatase, aminopeptidase) have also been observed.50 In addition to enzymic hydrolytic action, the local infiltration of white cells (monocytes and macrophages) augments degradative action, not least through the release of free radicals including peroxide.51,52 The latter is a general part of any non-specific inflammatory tissue response, but has long-term consequences for the integrity of any polymeric component of a glucose sensor. In the intermediate phase of the tissue response, the metabolic activity of inflammatory cells53 leads to a general modification of local tissue metabolic profile, and there are some apparent interspecies differences in this,54 which would make it difficult to extrapolate into man. Much depends also upon the size of the sensor burden upon the local tissue. For example, on one larger scale model system where cellulose sponge was implanted subcutaneously,55 extreme lowering of wound oxygen was observed at five days, as well as sequential variation of pH and CO2. The above changes not only feed back into further tissue reactive responses, but also lead to further development of unpredictable `solution' environments for a sensor. Thus, oxygen and lactate have been suggested as regulators of the wound healing process. High levels of lactate accumulation in wounds may improve wound healing through better collagen deposition, though this is contrary to the requirement for a minimal fibrous tissue in order to achieve undistorted tissue glucose monitoring.56 Quite apart from any physical cell/colloid accumulation, tissue remodelling and local fluid volume change, local lymphatics and capillaries are bioreactive components of tissue. During the wound healing process, they are known to undergo fluctuations in size, fluid throughput and perivascular organisation,

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inevitably impacting on fluid exchanges and solute filtration. These changes may also lead to capillary blood glucose supply variations and, therefore, unpredictable fluctuation in measured levels. Such physiological changes in blood flow were proposed as a reason for fluctuating patterns in oxygen response at a (cerebral) tissue implanted oxygen electrode developed by Clark et al.57 Non-traumatic implantation of a glucose sensor is not feasible currently, but clearly minimal trauma induction is a rational approach, and only truly achievable with a miniature device. Indeed, if dimensions can be made much smaller than those currently available, there would be major gains in reliability. However, as with conventional biomaterials, moving into an era of active control of the local tissue response through a local drug delivery strategy and device-incorporated bioactive agents may also change the outcomes. Thus, it is likely that implantable sensors will, in the future, use bioactive components for conventional implants. This would complement ongoing efforts aimed at the refinement of the bioelectrochemical transduction process itself.

5.4

Materials interfacing strategy

Direct biosample interfacing for a bioelectrochemical device is not feasible. Not only does glucose/O2 permeation into the device require better management in the absence of sample preparation, but also an interfacing material has to be provided to protect the internal components. As a consequence, selective encapsulation is mandatory. Therefore, specific investigation is required of membrane, interfacing and packaging parameters, inclusive of solute (e.g. glucose/oxygen) partitioning and transportation requirements.

5.4.1 Membranes for biosensor interfacing Polymers, rather than inorganic membranes, have been used from the start of biosensor technology to separate sample from the sensing components. As packaging and separation phases, they do offer some useful solutions to the interfacing of biosensors with complex sample matrices. This sub-area amounts to a materials sciences effort in the development of both low fouling and high selectivity barrier structures. Membranes may be classified according to their polymeric constituents (charged, neutral, amphiphilic), and their structural anisotropy. The most convenient and relevant for biosensors is the classification based upon available pore size. On the large scale, there are standard filters of 100 m pore diameter; microfiltration membranes of 1±10 m pore size are used to separate viruses and whole cells; ultrafiltration >10 nm membranes are used to discriminate macromolecules and colloids; reverse osmosis membranes of lower nominal pore size are able to resolve small organics and ions, though in reality these have no discrete pore architecture.

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In the case of porous structures, pore size variation allows for aperture control, and therefore for control over solute access to and from the device. It is likely that commonly used polyurethane external membranes, for implantable (needle) biosensors, operate as glucose diffusion controlling phases through their inclusion of a defined bulk phase porosity;58 certainly the polymer itself is impermeable. As a consequence, membrane technology offers a generic, adaptive platform for the design of a presenting sensor surface. Almost invariably, polymer membranes are used, and variously optimised according to the transport needs of charged, neutral and amphiphilic constituents.59 However, a wide array of chemical and structural features determine eventual interactions, including charge, surface energy, topography and the presence of specific functional groups. Individual examples of improving compatibility include the use of diamond-like carbon (DLC) coatings,60 as non-plasticised, robust, noncrystalline amorphous layers. Here, controlled analyte transport through DLC thickness variation was achieved, with the degree of surface fouling modified through surface profile and surface energy changes. In another example, surfactant incorporation61 into polymeric phases reduced fouling through creation of a quasi-fluid interface with high surface molecular mobility giving a mimic to natural soft material surfaces. The most striking form of biomimicry has involved direct adaptation of natural cell lipid membrane motifs and molecular components. These were embodied in particular in the phosphorylcholine zwitterionic layers pioneered by Nakabayashi and Chapman respectively,62,63 and are modelled on the external surface of the cell plasma membrane. Though these are functionally relatively effective, with phosphorylcholine being the key component, mechanical stability, however, requires covalent attachment to polymeric materials, detracting from the original fluidity properties of the natural plasma membrane. Also, the need for solute (e.g. glucose) permeability limitation through the lipid construct precludes the use of a high phosphorylcholine content. The fabrication of membranes for biosensors encounters various problems, particularly so when it is reproducible and well-defined porosity that is required. The operational problem is due to the presentation of a porous morphology to the sample matrix and thereby due to the tendency to take up protein aggregates into the pores, leading to progressive blockage. Our understanding of the dynamics of the pore blockage in relation to pore size and geometry is developing64,65 but remains incomplete. Progress in this area would be achieved by the rigorous definition of the term `porosity', a parameter unfortunately dependent upon the method used to determine it.66 There is also the possibility of membrane voids permitting antibody/ signalling molecule access to the antigenic constituents in the device, notably the enzyme or any background inert protein (e.g. albumin) to immobilise the enzyme. Moreover, avoidance of any protein release from the sensor demands

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robust chemical immobilisation, but even release of limited quantities of protein if they occur over extended time periods might lead to body sensitisation in susceptible patients. Paradoxically, release of soluble proteins into the body is more likely to be antigenic than that of larger protein aggregations made up of, say, crosslinked material. The membrane interface is thus an important safety feature for a bioelectrochemical sensor. Given the potential dangers of prions and the more classical hazards of viral particles, biocomponent origin, traceability and quality control will be as important to consider as the external membrane barrier in any implantable device.

5.4.2 Membrane property requirements Important surface properties to consider (Table 5.1) depend upon the biomatrix application. Whilst great emphasis was placed in the past on hydrophobicity and a negative charge, in order to nominally reduce the deposition of negatively charged cells and to avoid polar purchase sites for hydrophilic protein domain interactions, in practice, neither approach can be relied upon in most cases. Adverse surface interactions have proven rather more subtle. There is a strong possibility that membrane surface feature size and profile is important, and that an ideally `perfect', smooth surface would encounter minimum biofouling. Table 5.1 Key surface properties of polymeric membranes Chemistry Mobility Topography

Polar, charged, H-bonded, hydrophobic, ionisable Backbone, sidechain, plasticiser mobility Roughness (molecular, cellular), irregular/regular pattern porosity, pinholes

Also surface fluidity on the micron scale, analogous to cell membrane fluidity, may be a means of avoiding protein cell adsorption. Phosphorylcholine (PC), which has both positive and negative charges, is found on the external surface of red cells. It offers an elegant biomimetic solution in order to achieve low fouling.67,68 However, a comprehensive PC layer is likely to inhibit the transport of polar diffusible species and target macromolecules to the sensing element of a biosensor.

5.4.3 Conferred functional advantages of membranes Membrane and thin film technology may well in future make the difference between commercially viable and non-viable devices, viz. biochips. Thus, they confer a variety of benefits, which reduce the high level intrinsic specifications demanded of microfabricated biochips, so serving as a useful complementary

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Table 5.2 Benefits offered to biosensors by use of membrane technologies Membrane technology Surface modification Aperture control Chemical `gate' Sample clean-up Interferent control In-vivo biocompatibility

# Without surface modifications biochips are vulnerable

technology (Table 5.2). Surface modification allows adaptation of the base materials used for biochip production. Aperture control through pore size management allows for reduction of analyte access to the device and therefore, without sample dilution, enables the device to operate at concentration ranges where the biocomponent binding would normally be saturated. Thus, at high concentrations, bioaffinity is zero order with respect to concentration and, in that case, the biosensor becomes of little use. If the membrane has surface-attached pendent groups, then solute transfer through its pores will be affected by solute/ wall interactions. The greater the interaction, the lower the transport. Such chemical `gating', whilst not often used, could resolve mixtures of similar solutes and so underpin resolution by biochip arrays. Sample clean-up may similarly be achieved. There are significant advantages resulting from allowing biosensor detection reactions to be fully partitioned from the sample itself. For interferent control, a homogeneous or reverse osmosis type membrane is generally used since, in that case, microsolutes need to be rejected.

5.5

Membrane systems used in biosensors

Considerable versatility and ingenuity is required to translate membrane fabrication techniques at the macro-scale to the microfabrication of biosensors. Our past work has emphasised the former and, in general, solvent casting of preformed polymers has been used to coat planar and needle-shaped electrodes. Such techniques could be utilised in order to lay down conventional thick films (Table 5.3) for biochip manufacture, but more attention needs to be paid to edge effects and the ambient solvent evaporation conditions encountered when forming microsurfaces. More amenable to microscale deposition is in-situ monomer electropolymerisation. Whilst a conducting surface is obligatory for such a coating process,

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Surfaces and interfaces for biomaterials Table 5.3 Membrane materials usable for biosensors Membrane types PVC PVC-lipid PVC-surfactant Diamond-like carbon Cellulose acetate Polyurethane

Silane Sulphonated-PEES Polycarbonate Electropolymerised films Bi-layer lipid membranes

deposition can be highly controlled and solvent-related edge effects eliminated. Electropolymerised films, if conducting, can provide a further means of interrogating the biolayer if biolayer binding to a target is associated with charge related or conductivity behaviour modification of the film.

5.5.1 Thick membrane films at electrochemical sensors Microporous membranes Microporous membranes, based on polycarbonate through which linear track micron diameter pores have been etched, are in common use for particle separation. They are an ideal non-swelling structure for partial colloid and cell separation. In the context of biosensors, the pore size (and pore density) may be independently controlled. As a consequence, so is the aperture for driving down transport of a diffusible target solute into the biosensor. For example, in the case of oxidase-based biosensors, Substrate ‡ O2 ! product ‡ H2 O2

5:1

such as those for glucose and lactate monitoring, the low Michaelis constant (Km) is in this way extended beyond the upper concentration range of the fluid to be analysed, O2 demand is reduced and, as a consequence, an effective linear dynamic biosensor range may be engineered. The low permeability furthermore eliminates bulk convective flow influences, so that a sample viscosity and stirring independent system results.69 Microporous polycarbonate may be further modified with an outer coating of diamond-like carbon (an amorphous H, C alloy without a crystalline domain). The diamond-like carbon nanofilm has high haemocompatibility70 and, because it is deposited from a plasma source, exquisite control over coating thickness and therefore membrane aperture can be achieved. Silicones deposited on a microporous membrane71 also allow for controlled pore aperture but, because the silicone is O2 permeable and impermeable to organics, a larger aperture for O2 is afforded than, say, for a metabolite substrate such as glucose, pyruvate or oxalate. This allows oxidase-activated biosensors to be produced with extended linear responses, since the reaction (eqn 5.1) is then

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even less oxygen demanding. Using silane films, variously produced from the crosslinking of di- or tri- halogen substituted silane monomers, also gives a better haemocompatibility than the base material. A planar microelectrode array utilising the oxidase route for measurement could, in principle, be covered with a single microporous membrane and variously coated with modifying films of silicone or diamond-like carbon. This would create a family of sensors with a wide spectrum of sensitivities (calibration slopes) and linear limits that would allow for multiple redundancy, statistical data acquisition, drift free measurement and higher data security. The challenge is for membrane coating techniques to be carried out on a length scale that matches biochip fabrication. Homogeneous membranes for transport control CuprophanÕ is a well-known and well-established regenerated cellulosic membrane used for medical haemodialysis. It is crystalloid-permeable, colloidrejecting, and non-toxic. As such, it has been used in many reports in the past as a covering barrier layer for enzyme-based biosensors. However, its permeability and selectivity cannot be readily altered. On the other hand, PVC is, at first examination, an unpromising material for selective separation or controlled dialytic transport; it is organic solute

5.4 Response to lactate of a lactate oxidase based sensor that has an outer membrane of PVC permeabilised with the surfactant Triton X-100. As the weight-percentage of surfactant decreases the effective linear response and dynamic range of the sensor is extended.

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5.5 The cationic surfactant (MTAC) is incorporated into cellulose acetate (CA) as an outer membrane coating for pyruvate. As the surfactant content increased the effective response of the sensor also increased due to increased retention of co-factor within the proximity of the working electrode.

impermeable. However, it is an excellent support matrix for surfactant.72,73 Figure 5.4 shows the neutral surfactant Triton X-100 physically supported and entrapped in PVC at varying weight ratios and used as an outer barrier membrane for a lactate oxidase enzyme electrode. The lowering of surfactant content reduces an already remarkably high lactate ion permeability to optimise both biosensor sensitivity and the effective Km (linear range) of the integral oxidase of the biosensor. The addition of surfactants to a polymer membrane helps separate polymer chains in order to create voids for increased solute transport. In Fig. 5.5, cellulose acetate was mixed with a methyltrialkyl cationic quaternary ammonium ion surfactant (MTAC), which allowed better transport of anionic pyruvate across the resulting membrane. Importantly, the membrane cation was able to retain positively charged cofactors for the enzyme. As a consequence, only a relatively low drift was observed, even in cofactor-free sample. Porous membranes for lateral transport control The sandwich immunoassay format (antigen capture by antibody, identification via second labelled antibody) is an advantage for electrochemical detection of the binding event. It is necessary to separate attached second antibody from that which remains free in solution. However, in the case of the electrode being an integral part of the flow system, as is shown in Fig. 5.6 (using a system based on a silicon flow channel), the electrode element needs either to be subjacent to the capture antibody, or strategically placed further downstream. Transport of

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5.6 Sandwich immunoassay format within silicon microchannels where the detector electrode is either positioned at or downstream of the immobilised capture antibody site.

reporter antibody in a single step in an open dipstick planar configuration is more difficult, but has the advantage of being pumpless (Fig. 5.7), through exploitation of capillarity. The support material for such transport can be a microporous membrane. However, when nitrocellulose is used, it strongly adsorbs the second antibody (here with an alkaline phosphatase label) and lateral transport is therefore prevented. Nevertheless, the problem can be overcome by preadsorption of a blocking protein (Fig. 5.8). Unfortunately, this strategy cannot be used with cationic nylon for example (Fig. 5.9). As a conclusion, membrane materials have the potential to accommodate fluid microflows and, with judicial choice of the carrier surface (as with surface fouling), the affinity of mobile macromolecules can be modified and their transport behaviour controlled.

5.5.2 Non-conducting electropolymerised films Electropolymerisation of ultrathin films on working electrode materials is a means of film coating that is used with equal facility for macro-scale and microscale electrodes. The basic electrode surface develops a protective layer analogous to that of films used for corrosion protection. One functional aim is to produce low molecular weight cut-off films in order to exhibit high H2O2

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5.7 Open flow planar membrane format in which capture antibody is immobilised on solid phase support; sample and reporter antibody are applied from bulk solution.

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5.8 Alkaline phosphatase labelled antibody transported laterally using albumin preblocking of the carrier surface.

selectivity, which is the primary product for biosensing of oxidase-catalysed reactions (eqn 5.1). Such films may also improve haemocompatibility, extend the linear range for the analyte of interest and help better understand the interactions between sample and biosensor surface. Polyphenolic films: biocompatibility and selectivity The electrooxidation of phenol in solution produces a half wave at about +0.65V vs. Ag/AgCl. No reduction wave can be observed on reversing the voltage sweep in cyclic voltammetry (Fig. 5.10). This is a strong evidence of a stable, inert film on the electrode surface, which accumulates on multiple sweeps until a limiting

5.9 Albumin blocking of nylon does not overcome strong antibody adsorption.

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5.10 Electropolymerisation of phenol (10 mM in PBS pH 7.4) on Pt working electrode. Scan rate 50 mV/s. Following initial oxidative wave at +0.65 V vs. Ag/AgCl due to electrooxidation of monomeric phenol, polymerisation occurs such that Pt becomes insulated preventing further oxidation of phenol. The process is therefore self limiting.

5.11 Effect of polyphenol modification on selectivity profile of Pt.

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insulating layer is formed, which precludes monophenol oxidation. This film layer provides an entirely different selectivity profile for the device (Fig. 5.11), which in particular is seen to be relatively accessible to H2O2, whilst largely eliminating transport of common blood interferents. Catechol, in this case, is not considered as a constituent of blood but as a potential end product of a redox indicator reaction cascade usable for monitoring dehydrogenase reactions.74,75 The haemocompatibility of a surface is usually improved by polyphenol. Although some drift is observed, the comparison with the bare electrode is still highly favourable (Fig. 5.12). These films may also offer some advantages not

5.12 (a) Susceptibility of bare Pt working electrode to whole blood exposure; (b) Stabilisation of Pt working electrode to whole blood exposure through polyphenol modification.

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only for microfabricated structures used in blood but also in protein loaded solutions as they might be considered for proteomic arrays. Polyphenol variants Phenolics with more complex structures and pendent groups would be expected to modify the selectivity and biocompatibility performances. A range of these have been tested (Fig. 5.13). Of these, poly(rosolic acid) appeared to be the most stable (Fig. 5.14). There is further possible manipulation of the interface using different conditions. For example, polymerisation appears to be faster at higher potentials and the net charge requirement becomes lower. These are free radical propagated polymerisations following the initial phenol oxidation therefore forming efficiencies might be expected to be different. Higher solution pH gives more porous films and selectivity is due to a cumulative effect depending on deposition times. Consequently, individual films may be tailored or used in systematically varied biochip arrays.

5.13 Select ion of phenolic monomer s used during format ion of electropolymerised coatings on Pt working electrodes.

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5.14 Selectivity of Pt modified with poly(rosolic acid) coating following exposure to whole blood.

Polyphenols with surfactant entrapment Polyphenols provide entrapment of `bystander' solutes, in solution, provided that these are of sufficiently high molecular weight. Surfactants are of interest, since not only are they sufficiently large to be physically entrapped in the films, but their solubilising effects and variable charge properties (Table 5.4) would help confer a modified function to the films, with possibly better haemocompatibility. Figure 5.15 suggests that fouling in blood may be reduced over time. In all instances, with film coatings, the level of stability in blood is excellent compared to that obtained with a bare electrode (Fig. 5.16). These data strongly indicate that the base Pt (and other) electrode materials are not sufficiently reliable for use in biofluids, no matter how sophisticated the Table 5.4 Surfactants used for entrapment in polyphenols Surfactant

Polarity

Formula weight

Critical micelle concentration (cmc) (mg/ml)

Non-ionic

8400

100

Adogen 464 (methyltrialkyl (C6±C10) ammonium chloride)

Cationic

368±452

n/a

Taurocholic acid

Anionic

537.7

6.7

Pluronic F68

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5.15 Effects of incorporating surfactant within poly(rosolic acid) film on stability of Pt electrode following exposure to whole blood.

5.16 Reduction in sensitivity of bare Pt following exposure to whole blood, shown as normalised response.

electrode may be (e.g. a biochip array), unless some specific means of protecting against fouling has been used.

5.5.3 Conducting polymer films Planar conducting films, whilst potentially usable as a membrane barrier or as a biocompatibility inducing interface, can also operate as reporters for bioaffinity

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5.17 Species used for the formation of conducting polymer systems.

reactions. Moreover, they can be formed over microelectrodes with the ease of non-conducting phenolics.76 A range of organic compounds have been used to form conducting films (Fig. 5.17), but polypyrrole appeared to be especially advantageous. Conducting polymers can be switched between the conducting and insulating states. Film conductivity correlates with the prevalence of mobile cationic species (polarons/bipolarons) available on the polymer backbone.77 Polaronic conductivity is due to a delocalised positive charge spread over four pyrrole units, whereas electronic conductivity involves the conduction band of the polymer. The films generally used in aqueous solution demand relatively low potentials for redox switching and changes of the state of conductivity. Film deposition is initiated by the oxidation of a monomer to produce a radical cation, which then reacts with other monomers or monomer radicals to go to a dimer stage and then a trimer stage and so on. Oligomer formation then moves on to nucleation and polymer elongation. Importantly, the charged polymer incorporates solution anions. If these anions are bioreceptors, then a biosensor results. The bioreceptor may be co-entrapped in the films without formal charge-based incorporation, provided that a simple anion is present in solution to facilitate the polymerisation process. Inorganic anions as well as anionic oligostructures, such as sodium dodecyl sulphate (SDS) and toluene sulphonate, can be entrapped as stabilising counter ions.

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Conducting polymer electropolymerisation A planar two-electrode arrangement (Fig. 5.18) comprising interdigitated gold (Au) electrodes if used as a polarisable electrode can allow surface deposition of polypyrrole. Once nucleation has occurred, film growth can be followed during cyclic voltammetry (Fig. 5.19). The counterion incorporated affects the voltammetry signatures of the formed films, which is observed when studying the reduction peaks (Fig. 5.20). However, avidin incorporated into a film, when

5.18 Electrode arrangement for planar film impedance spectroscopy. A schematic indicates the 15 micron gap between electrode fingers.

5.19 Formation of polypyrrole as monitored by cyclic voltammetry. There is a continued increase in the magnitude of the oxidative wave with subsequent potential scans.

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5.20 Effect of incorporated counter ion on the reduction peak signature of poly(pyrrole).

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used as a receptor for D-biotin, shows little effect on the cyclic voltammogram. This would seem to indicate that voltammetry is not an appropriate means of detecting receptor changes in the polypyrrole. Some alternative methods of analysis are therefore required. Interdigitated structures (Fig. 5.18) can be bridged by means of polypyrrole. Bridging can then be used to measure film conductivity with any given counterion. To achieve this, a low ionic strength solution is needed to maximise uptake of a receptor molecule and to control the extent of bridging achieved between electrodes. A relatively high polypyrrole concentration is also needed to avoid inhibition of film growth. It is important to recognise that film structure may vary depending on the counterion used. SDS leads to confluent growth and avidin is associated with discoid, multipoint growth. Such variation may have an effect on analyte access to the film, available surface area for binding and the degree to which non-specific surface binding of proteins and cells occurs at the sample interface. Impedance spectroscopy at planar polypyrrole films Once formed on a surface the polypyrrole film can be switched from an insulating to a conducting state based on the impressed polarising voltage. However, a particular redox state can also be sustained, determined by some intermediate polarising voltage. The set redox state can then be followed by twoelectrode electrochemical impedance spectroscopy (EIS). Two electrode EIS can be achieved using a 20 mV RMS sinusoidal potential, and an impedimetric spectrum taken between an extensive frequency range, from 5 Hz to 13 MHz. Independent measures of polaronic and electronic conduction can be obtained, since they take place on quite different voltage oscillatory time scales. A Bode plot is a useful graphical representation of impedance data, combining capacitive current measurement (out of phase with sinusoidal voltage) with impedance measurement (in phase with sinusoidal voltage). Impedance (Z) defines the relationship between applied potential and current and, more 0 formally, is a vector quantity composed of a real (in phase Z ) and an imaginary 00 0 00 Z ‡ iZ where `i' is the (out of phase Z ) part related by the equation Zpˆ complex number and is defined by i ˆ …ÿ1† and the modulus 0 00 jZj ˆ …Z 2 ‡ Z 2 †0:5 . Physical interpretation of the data is problematic. However, the observed data can be modelled using derived equivalent circuits comprising arrangements of resistors, inductors and capacitors. In practical terms, an electroinactive product or outcome of a bioaffinity reaction can be registered and a two-electrode arrangement is usable, provided that the redox state of the film remains stable. Real films show distinct separate zones respectively of electronic and polaronic conduction (Fig. 5.21). In ionic solutions, such as a phosphate buffer solution, polaronic conduction is augmented because the presence of the mobile anions in the film facilitates charged-linked electron transfer.

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5.21 Two electrode impedance spectra at IDEs.

Impedimetric measurements of LH and biotins For a model system incorporating antibody to lutenising hormone (LH), a polaronic phase is identifiable (Fig. 5.22) but, more importantly, there is a capacitance reduction and an impedance increase after binding to LH. This is seen only when a redox cycle is imposed after the binding took place. This provides an indication not only of how antibody can be entrapped within

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5.22 Reagentless binding with LH ± upper capacitance peak, and lower impedance curve trace the zero LH response. LH concentration 100 IU/l.

polypyrrole, while remaining active and accessible to a peptide such as LH, but also of how, in the absence of any apparent voltammetric changes, the impedimetric signature is altered to thereby achieve reagentless binding recognition. The inherent principle should be translatable to other affinity pairs. Thus, for avidin entrapped in polypyrrole, responses are seen for biotin and biotin analogues that are significantly different from the entrapped urease control with regard to phase angle (Fig. 5.23(a)). The biotin derivatives used are amidocaproate succinimide ester (biotin ester) and biotin amidocaproate 3sulpho ester (biotin sulpho). There is firstly an increase in phase angle observed when exposed to phosphate buffer (this is a background effect that would need to be taken into account even if practical measurements are carried out), but the effect of biotin binding is clear-cut. A key feature of all observed binding-induced phase angle changes is the need for one preliminary redox cycling step to unmask the effect on capacitive behaviour (phase angle). Possibly such cycling allows for realignment of polymer chains around the newly formed complex, in turn changing the polymer conformation and, therefore, the ring alignment and polaronic conduction. The importance of the original gold-polypyrrole surface in conditioning the response is illustrated (Fig. 5.23(b)) in the lack of significant biotin response where an electrode had been reused following formic acid cleaning to remove previous films. The subtlety of this effect is underlined by the finding of basic Bode plot profiles for recycled electrodes that are identical to unused electrodes. Thus, either the gold electrode interface is directly important for conditioning cross-IDE impedance, or the gold surface affects the superimposed film structure. The avidin biotin complex is a useful model to study, but the high binding affinity of avidin (KD ˆ 10ÿ15 L/M) does not readily allow concentration

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5.23 Effect of base IDE condition on polypyrrole film formation. (a) unused electrode (b) recycled electrode.

dependent effects to be followed. These, however, can be seen for a LHantibody combination but it is also notable that with the polyclonal antibody preparation used, notwithstanding a high LH affinity, there was little response in contrast to a monoclonal antibody. Possible reasons for this include blocking contaminants in the polyclonal preparation, molecular heterogeneity, distorting polymer conformation changes following binding and antibody aggregates, reducing microenvironmental effects within the bulk of the film structure. The monoclonal-LH system demonstrated no significant response in phosphate buffer, but response to LH was attenuated by pre-incubation of the LH with bovine serum albumin (3 g/L), tested up to 800 IU/L LH. The albumin possibly blocked binding or there may have been bulk solution associated effects with LH that prevented antibody interaction. Whatever the reason, such complex macromoleculer interactions highlight a further challenge to measurements in biological samples. This is in marked contrast to the use of impedimetric measurements at conducting polymers for gas phase affinity sensing, as developed for artificial nose technology,78 where interferences are minimal.

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5.6

Microflows as surrogate, renewable barrier films

5.6.1 Needle biosensors and in-vivo monitoring Notwithstanding the structural elegance and geometrical precision of MEMs based sensors for in-vitro use, in-vivo insertion, even for percutaneous blood sampling and intravenous lines still relies upon needle technology. The fundamental point is that the cylindrical geometry, miniaturisation and high mechanical specification of needles satisfy the dominant issues for healthcare use and take precedence over the structural and length scale elegance of MEMS devices. In the case of silicon-fabricated devices, there is a further concern over long-term implantation, in that dissolution of silicon and local elemental release may be of significance, an issue that needs quantitative consideration if, say, porous silicon is to be used for in-vivo device construction. The needle design has been consistently used for implantable enzyme-based biosensors. Inevitably the greatest attention has been paid to glucose sensing in view of the healthcare importance of diabetes. One classic construction route for this has used two functional layers aside from the chemically crosslinked enzyme layer of glucose oxidase.69 The first layer is a selective low molecular weight cut off barrier based on a sulphonated polyethersulphonepolyethersulphone (SPEES-PES). This robust ionomeric barrier mostly excludes ascorbate, urate and acetaminophen (TylenolÕ) as major electrochemically active interferents on the combined basis of size and charge. The degree of sulphonation is important and may explain why a high selectivity has not been seen by some.79 The outer layer covering the antigenic enzyme layer must interface with the biofluid. It is also a multilayer porous polyurethane serving to reduce glucose transport to the enzyme layer, retain O2 transport (eqn 5.1) whilst providing mechanical integrity and high resistance to the degradative action of free radicals potentially generated during chemical reactions. The laminate formulation used by us is a combination of five pre-polymer polyurethane layers and a single carbonate polyurethane (CorethaneÕ) outermost layer known for its tissue compatibility and low degradation rates.

5.6.2 Open microflow for in-vivo monitoring Needle glucose electrodes used in tissue for short-term (~24 h) monitoring should, in principle, be acceptable for diabetic home use. In practice though, percutaneous, temporary insertion for such minimally invasive monitoring is difficult to achieve. No matter how stable such devices are in vitro, even in blood, the tissue response to the implant, however, demands firstly a stabilisation period of some hours and then in-vivo calibration in the tissue. Indeed, tissue glucose values, both at steady state and under dynamic conditions, show a clinically acceptable match for nearly all current devices. The reasons for this are complex and reflect transcapillary transport lag time, the limit of the

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microcirculation to supply and remove glucose, the nature of the inflammatory exudative tissue response and the distance between the capillary bed and the implanted sensing surface. The last is compounded by the barrier function of the interstitial tissue `mesh'. Open microflow (Fig. 5.24) allows for delivery to the tissue implant site of a film of hydration fluid delivered over the sensing surface of the device and then into the tissue itself. Because subcutaneous tissue mostly operates at a negative hydrostatic pressure, fluid transport does not necessarily require physical pumping and at a net outflow diameter of 0.2 mm, 1±2 L of buffer enters the tissue. The amounts are therefore extremely low, and well within the range of volume exchanges that occur at the capillary bed under physiological conditions, as driven by net positive and negative transcapillary pressures in the Starling mechanism. The externally introduced fluid flow creates a mobile film over the sensing surface that helps to reduce colloid/cell access to the sensor and reduces fouling. Equally important, the juxtaposed tissue is hydrated, and such hydration of connective tissue reduces the diffusional barrier presented to microsolutes. Hydraulic permeability of oedema tissue, for example, is known to be increased several orders of magnitude in comparison with normal connective tissue. The practical consequence of open microflow is that sensor `run in' time is reduced from a typical 3 h down to a mere 30 min. Also the tissue:blood mismatch is reduced as is the commonly observed temporal lag (up to 12 min) for tissue glucose changes. These advantages are most in evidence in the animal model where tissue-related monitoring error appears all but eliminated.

5.24 Percutaneously implanted needle enzyme electrode showing hydration zone at needle tip following pumpless flow of fluid into the dermis.

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Translation into humans shows promise, however, microfluidic design changes are a likely requirement given the interspecies differences in connective tissue composition and capillary bed behaviour.

5.7

Microfluidics and biosensors

While used as `dip in and register' devices, biosensors hold the promise of deskilled analysis. One element of their current attractiveness consists in the fact that they can be produced by microfabrication techniques, and are therefore ready for miniaturisation at single device or industrial scale production. Thus, regardless of this, achieving full analytical benefit depends upon sample presentation. Thus, sample flow, dilution, background composition and the presence of interfering substances all add to the so-called matrix effect in biological samples, and all variously conspire to distort the basic signal readout. It is a fact that sample presentation and delivery have been a much neglected area of research on biosensors. This is also one reason responsible for their lack of widespread uptake, either for extra-laboratory testing, or in drug discovery arrays. However, thanks to rapid advances in microelectromechanical systems (MEMs), it is now possible to design and manufacture high precision flow structures,80 predominantly in silicon or glass. By using these structures, researchers have been able to create designer fluid flows that offer the prospects of precision fluid delivery to targeted locations via valve controllable flow paths. Beyond any size scaling effect, however, fluids under microfluidic flow conditions exhibit quite distinct fluid dynamics. Here, viscous forces are dominant, and any irregularity or turbulence in the flow pattern evens out in order to give a laminar flow, with precise, quantifiable solution transport profiles through an entire length of the microchannel network regardless of channel interconnections and overall geometry. Microchannels or capillaries (usually 0.02 to 2 mm diameter) can be regarded as reaction microenvironments. They exhibit two very specific characteristics. First, they present small internal volumes that can be easily and rapidly filled by fluids. Second, fluids moving inside these channels remain laminar even at very low velocities, with no evidence of turbulence. A microchannel thus constitutes a special delivery vehicle for self-contained and finely controllable sample exposure of a strategically positioned biosensing surface along the flow path. Moreover, the use of miniaturised flow cells giving Reynolds numbers of less than 2000 (vide infra) permits non-turbulent fluid flow so independent, parallel flow streams can be accommodated without turbulent mixing.81 A microfluidic channel presents a relatively high surface area to the bulk sample volume, which thereby enables a wall-immobilised reagent to be highly effective in reacting with bulk solute constituents, e.g., through binding and catalysis. These high surface to bulk interactions also facilitate chromatographic principles of sample separation.

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Laminar flow allows delivery of chemical species inside capillaries with remarkable precision. The application of different streams of a multiphase, laminar flow system also allows a reaction to take place exclusively at certain regions of the flow system, either between a stream and channel wall or between two adjacent streams. As a consequence, such devices could be potentially used to carry out in-situ microfabrication experiments, thus paving the way to innovative micro and even nanotechnology strategies.

5.7.1 The Reynolds number In 1883, Osborne Reynolds demonstrated that, under specified conditions, stable laminar flow switched to unstable turbulent mode. He introduced a fundamental equation, which gave a range of dimensionless numbers, the Reynolds numbers, defining the transition from laminar to turbulent flow. The Reynolds number is defined as the ratio of inertial vs. viscous forces and is used in momentum, heat, and mass transfer to account for dynamic similarity It is calculated using: Re ˆ

D
V
 

5:2

where Re ˆ Reynolds number, D ˆ characteristic length, V ˆ velocity,  ˆ density and  ˆ viscosity. Due to the large number of parameters, there is a great degree of flexibility available regarding the flow cell construction and arrangement. Typically, a single flow within a straight pipe with smooth walls and Reynolds numbers below 2000 is both stable and laminar. Between 2000 and 2300, a transitional flow is observed. Above Reynolds numbers of 2300, the flow becomes turbulent and unstable. Most microfabricated fluid systems operate under low Reynolds number conditions. Since in small channels only small liquid masses are moved and flow velocities are typically low, this ratio is generally small for any given viscosity. Under these conditions, an aqueous stream in a microchannel will behave similarly to viscous oil-like flow at the macroscopic level. Consequently, it is possible to flow two aqueous streams of similar viscosity side by side while preventing turbulent mixing.82 Nonetheless, if flow is subject to an abrupt change of channel geometry or encounters obstacles, it will change from a laminar to a turbulent regimen even at extremely low Reynolds numbers (Re40-300).81 Although turbulent mixing does not normally occur at low Reynolds numbers, diffusion still occurs, and has been utilised by Manz and co-workers to generate microfluidic mixing, otherwise difficult to obtain at such low Reynolds numbers.83 A key consideration when using microfluidics is to understand and control flow characteristics (velocity profile, shear force, etc.) which is a complex area of fluid dynamics that relies heavily on computational modelling in order to

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predict these outcomes. The Navier-Stokes equation is often used to characterise fluid flows provided that the initial conditions are properly established.84 The equation is based on the balance of forces summarised as follows: Transient Convective Body Pressure ‡ ˆ ‡ inertial force inertial force force force

‡

Viscous force

5.3

However, in steady flow, the transient inertial force vanishes. Also, if the fluid is considered as ideal, the viscous force term disappears, simplifying analysis.85

5.7.2 Design and fabrication of microfluidic devices The main trend currently is to consider the biosensor as an integral part of a complete miniaturised analytical system rather than as a separate entity, thus the concept of a micro total analytical systems (TAS) involved microfluidics as part of the system.86,87 A scaled down fluid delivery system can be seen as equivalent to the distributed capillary networks of natural tissues, which permits conservation of sample volume and exquisite flow control and fast switching of flow paths and delay lines. All contribute to the desired, near-ideal conditions needed for the fine-tuning, multiplexing and efficient operation of bio- and other sensor systems. Microfluidics is essentially a basic, though less discussed, component of a TAS. Here, the manipulative steps of an analytical laboratory, notably sample separation, clean-up, reagent introduction, reaction profiling concentration optimisation and analytical endpoint detection are incorporated into a miniaturised structure, alternatively considered lab-on-a-chip. Miniaturisation of many of the relevant processing components has certainly been achieved, but none would be operational without the mobile element embodied in the integral microfluidics; much more needs to be done in developing the technology. Microfabrication of biosensors scales down reagents requirements, a major financial advantage, and also facilitates greater sample throughput. The method requirement is for microfluidics. Before fabricating a microfluidic system, time has to be spent in considering the precise application. Thus, the eventual operating conditions need to be specified. One important choice is that of the basic material (resin, polymer, metal, glass or silicon) from which the device will be fabricated. The material has to be chemically inert and resistant to the various solvents and solutes to which it is exposed during the operational lifetime. Each type of material is associated with different types of manufacturing methods and quite different manufacturing costs, which can be more or less important depending on the individual material used. When an aqueous phase is used, relatively inexpensive polymeric materials like poly(methyl methacrylate) (PMMA) and poly(dimethylsiloxane) (PDMS) can be used to fabricate flow-cells. However, some solvents will alter if not

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completely erode these polymeric materials. As a consequence, for a microfluidic device handling a wide variety of solvents (aqueous to organic, polar to apolar), it becomes vital to choose more expensive materials that are chemically inert vis-aÁ-vis these solvents, as well as the solutes they will be carrying. Such materials (metal, glass, silicon) require more complicated, and therefore more expensive, techniques for etching and assembly, reversibly where possible, into a well-sealed microfluidic enclosure. It is also crucial in this instance to choose the right component to dispense and transport the fluids from a reservoir into the microfluidic device. Syringes have to be of glass and the tubing and connectors made of solvent-resistant but still flexible polymeric material, like PTFE (polytetrafluoroethylene) or PEEK (polyetheretherketone), as commonly used in liquid chromatography. Once the choice of material has been made, various etching/assembly techniques available need to be investigated. Microfabrication techniques, in their most basic form, rely on thin-film or thick-film metal pattern formation associated with photolithography.88 This technique has already proved highly efficient and successful in the mass production of silicon chips for the semiconductor industry. A thin metallic layer is deposited onto a substrate (silicon or glass), typically via spin-coating, after which a photoresist is added. Then, a patterned photomask is applied and the whole system irradiated by UV light. The unmasked portions of the photoresist are etched and the exposed metallic areas are then eroded either chemically in solution (wet etching) or by exposure to a plasma (reactive ion etching).89 The metal layer can be replaced by a layer of silicon oxide (SiO2) that is grown on a pure silicon substrate in a furnace over many days. The SiO2 layer can be etched using the same techniques as above. Silicon oxide and glass, which are similar, can be easily etched using HF (wet etching). The microfluidic channels formed can be sealed by anodic bonding of a glass slide or cover slip to the glass or silicon/silicon oxide substrate. The lift-off technique is very similar to the classical photolithography technique except that the photoresist is first deposited onto the substrate, then partially etched by UV light using the patterned photomask. The metal layer is then deposited on both the exposed substrate layer and the remaining photoresist. The photoresist is then dissolved (lift-off) leaving the patterned metallic layer. Anisotropic etching is capable of producing three-dimensional structures as opposed to photolithographic techniques. It is performed with heated alkaline solution using SiO2 or Si3N4 masks. The etching rate is highly controllable and stops at layers of impurities. It is dependent on the crystalline orientation of the substrate. In silicon etching, for example, the most commonly used etching agents are potassium hydroxide and tetramethylammonium.90 Powder blasting is a technique that can be used to generate microstructures in brittle substrates such as glass. It employs a masked substrate and an eroding

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beam of 30 micron sized alumina (or sand) particles in a powder form. This powder is propelled using a compressed air stream, which is able to etch channels down to 100 micron width and 20 micron depth.91,92 Many of the techniques employed for microfabrication are borrowed from the semiconductor industry. However, since silicon is hydrophobic by nature, it may not be suitable for biological studies, and the use of polymers represents a more attractive and cheaper proposition. Poly(methyl methacrylate) (PMMA), for example, can be used and is easily etched using standard mechanical tools or using laser ablation techniques to form very precise microchannels. Hot embossing is a technique capable of generating low-cost devices for the biotechnology industry. It utilises a hot plate with the substrate being pressed between the hot plate and another patterned metallic plate. Micro-injection moulding requires that the polymers used are first reduced to granules, melted above their Tg (glass transition temperature) and injected into a mould, generating microfluidic channels. Another polymer that can be used is Poly(dimethyl siloxane) (PDMS), a clear rubbery, crosslinked polymer that can be easily cast into a patterned metallic or silicon mould, thus generating microfluidic devices at very low cost, ideal for prototyping.93,94

5.7.3 Detection Chemical and biosensing devices integrated into fluidic arrays at multiple locations can allow sequential monitoring of reaction transients or of the kinetic development of reaction profiles along a given flow channel. Little additional structural distortion of precision-formed fluid channels results with wall embedded devices. More accurate kinetic measurements thereby achieved thus can add to end point estimations. In electrochemical sensors oxidisable or reducible molecules generate measurable current flow. With stable fluidics these current responses become further stabilised, and signal to noise is reduced, important for low current outputs. Neurotransmitters and the products of some enzyme reactions (oxidoreductases), either directly or via electron transporting indicators, have been measured in this way and even when using well-known chemistries there are additional benefits due to miniaturisation and the fine tuning of signal outputs. Through a combination of biosensor arrays and variable flow conditions, a spectrum of multiple responses can be obtained giving better sample profiling allowing for greater biosensor redundancy and ultimately improved data reliability. Sample concentration changes are also possible in miniaturised structures through controlled evaporative loss and, in the case of ionic solutes, adsorption/release from electrically addressable surfaces give additional routes to sample adjustment thus allowing the analytical `reach' of biosensors to be extended.

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Control over diffusive solute transport from the bulk sample to the biosensor surface can be augmented by in-situ polymerisation of thin conducting and nonconducting films (vide supra), readily formed within the confines of a closed microchannel. Such thin films do not distort flow, but do allow co-entrapment of biological or non-biological reagents for biosensor formation. Fluid conductivity measured with micro-electrode pairs across flow channels or specific ions measured with miniature ion-selective electrodes allow for both profiling of a sample and facilitate sample status monitoring in addition to any integrated biosensor for a specific target. A special opportunity arises through optical fibre or waveguide based optical interrogation of flow components, and detection of an intrinsic solute property (e.g. protein fluorescence) opens up the additional armoury of spectroscopy.

5.7.4 Diffusion and sample separation Laminar flow in microchannels, depending on the flow rate, may allow or prevent the diffusion of solutes from one stream to another. Diffusion becomes more prevalent as the flow rate decreases and the flow path increases. It is also possible to discriminate between analytes depending on their molecular weight. Low molecular weight compounds diffuse more readily than high molecular weight solutes for which diffusion can be almost completely prevented, reducing fouling of a sensing electrode surface. Figure 5.25 shows a schematic microfluidic device, where a bifurcation is used to converge fluid streams in a 200 m wide flow channel. The value of this flow phenomenon is that there is controlled laminar flow over any electrode located in the wall of the flow channel, so convective transport and therefore response of an electroactive species can be correctly modelled. Furthermore, a protective fluid film can be used over the intramural electrode to compartmentalise the electrochemical (or chemical) detection sequence from a potentially contaminating

5.25 Bifurcated flow cell schematic for non-mixing of parallel microflow.

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sample stream. The protective effect primarily applies to microsolute monitoring, where diffusion coefficients are at least one order of magnitude greater than for, say, surface-active proteins; cellular transport, moreover, across the two streams is virtually eliminated. Even with a protruding electrode it may be possible to have a viable monitoring set-up with the protruding device allowing for a high exposed microsurface area and therefore higher currents. In a study on the difference of diffusivity between a high molecular weight solute (glucose oxidase) and a small molecule hydrogen peroxide (H2O2), generated by the reaction between glucose and glucose oxidase, it was shown that with glucose oxidase in the stream adjacent to the electrode and glucose in the opposite stream, a reaction occurred at the interface between the two streams, generating H2O2. With increasing flow rate, the electrochemical current generated by hydrogen peroxide at the electrode decreased and a faster flow reduced the diffusion time for hydrogen peroxide. This behaviour is similar to that of a separation or filtration membrane but in a fluidic, thus renewable, form preventing biofouling of the sensitive sensing surface.

5.7.5 In-situ microfabrication Many experiments have been carried out using the unique behaviour of fluids inside microfluidic channels. Control over reagent delivery has been used to precipitate solid structures in flow channels such as linear wire electrodes or silver and conducting polymeric fibres.81,95 However, microfabrication of complete barrier structures, like free-standing ultra-thin membranes remains a challenging task. Thus, attempts at precipitating cellulose acetate from 0.2 wt% acetone solution at an aqueous stream using first a 1.5 mm (width)  0.5 mm (depth) channel and then a 1.2 mm (width)  0.6 mm (depth) channel led to aggregates of precipitated polymer but not membranes. The formation of distinct membranes with some mechanical strength was achieved by using a polymer that was not easily redissolved by the flow stream. Thus, interfacial polymerisation of nylon 6,6 was attempted. Adipoyl chloride (62.5 mM) in xylenes and 1,6-diaminohexane (62.5 mM) in distilled water were used. Interfacial polymerisation in the diffusion zone led to nylon 6,6 membrane instantly. A solid anchoring point proved to be a key requirement in order to achieve stable membrane attachment. Further development of permeable membrane structures should allow for an interfacial barrier for sensor protection, and could complement separation by liquid-liquid interfaces.

5.8

Conclusion

Though biochip-based and other precisely configured biosensors are precision structures with the opportunity of greater measurement reliability using multiple arrays, the need remains for reliable and stable interfacing. Relevant research in the

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areas of materials, membranes, fluidics and films is already available to help adapt and `refashion' biochip surfaces for practical use. A combined effort at this stage is more likely to help deliver viable products for the future, especially in healthcare. Admittedly the advantages may be of overwhelming importance only if measurement is required on a continuous basis or if unmodified (e.g. undiluted, non-deproteinised) samples are to be assayed. However, the over-the-counter medical diagnostic market is a prime example where unmodified samples will be measured, which is likely to benefit from biochip configurations of biosensors as truly reliable and therefore `smart' systems, provided that stability in the sample can be guaranteed. To this end, increasing interest in barrier membranes is anticipated. However, fluid interfaces and in-situ microdeposited membrane barriers could lead to major augmentation of microdevice performance with considerable economy of sample volume requirements.

5.9

Acknowledgements

The authors would like to thank EPSRC, JDFI, BDA and the EU for funding of the ongoing studies presented in this review.

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6

Micro- and nanoscale surface patterning techniques for localising biomolecules and cells: the essence of nanobiotechnology Z A D E M O V I C and P K I N G S H O T T , The Danish Polymer Centre, Denmark

6.1

Introduction

Most reactions in biology occur not in solution, but at interfaces, such as cell membranes. The need to improve the quality of our lives has led to substantial research and development in manipulating material surfaces in an attempt to mimic and control biological reactions that take place at interfaces. Medical implant materials, for example, have played a pivotal role in bringing surface concepts to biology. In the late 1940s the first biomaterials, as we known them today, were developed.1 Almost in parallel, scientists began studying the properties of biomaterial surfaces, which included investigations involving protein-surface and cell-surface interactions, and ways of modifying these surfaces in order to control so-called biointeractions. Present surface science has had considerable impact on biology and medicine. Implant biomaterials,2 affinity chromatography,3 gene chips4 and diagnostic arrays,5 cell culture surfaces6 and biosensors7 are examples where surface technologies have been applied to biological problems. These applications have been largely driven by an early appreciation of the fact that surfaces provide the platform for biological reactions to take place. An optimally functioning human being is made up of tissues that consist of highly organised structures of different cell types, each performing different but complementary functions. All the cells are surrounded by topographic and chemical cues that can align and guide cells.8,9,10 The specific protein-surface and cell-surface interactions are most likely to be strongly influenced by a combination of surface chemistry and surface topography,11 arranged, moreover, into well defined patterns on the surface at both lm and nm length scales.12 The relative importance of these two factors remains poorly understood because we are still developing techniques to create patterns and the analytical methods capable of actually visualising such features. In addition, results achieved so far have shown that topographical modifications are accompanied by chemical variation. Hence, controlling the topography and chemistry of polymer surfaces independently is an important issue. The effect of surface topography on cellular

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and tissue response to implanted biomaterials has been known for the past 20 years. Cells have been shown to respond to micro- and nanoscale structures and to topography. Slightly roughened surfaces show increased osteointegration, reduced fibrous encapsulation and enhanced integration of implant materials,13,14,15 a result of increased adhesion of connective tissue cells. An ability to properly engineer cell-surface interactions and to control the spatial location of cells is highly desirable, and essential for creating in-vitro mimetics of physiologically relevant interfaces. Understanding how the surface chemistry and structure of a material can be used to control the biological reactivity of a cell interacting with that surface is the ultimate challenge for biological surface science. Ideally, the surfaces will need to have well-defined arrays of biorecognition sites designed to react specifically with the cells, since many of the important functions of cells depend on the arrangement of biomolecules at their surfaces.16 A large number of fabrication tools exist for tailoring a biomaterial surface. Chemical surface modification can be combined with designed microstructures and nanostructures, aimed at matching biological components or for inducing a desired biological reaction. Controlled protein adsorption is crucial to optimise cellsurface interactions, since cells recognise the protein layer and not the underlying surface. Surfaces must be developed that control the conformation and orientation of proteins with precision so that the body will specifically recognise them. The simplest method for immobilising a biomolecule on a surface is physical adsorption, but this is a purely random event and can lead to denaturation and loss of protein function or displacement by less desirable proteins. A more stable kind of protein immobilisation involves creating covalent bonds between the protein and solid surface. In addition to efforts made to optimise and control protein orientation at surfaces, a new research direction in protein immobilisation is taking place, namely protein patterning.17 It can be defined as protein immobilisation within specific locations in either two- or three-dimensional space.18,19,20 To date, research in this field has had considerable success on the micrometre length scale with patterning in two dimensions, primarily with single proteins. The generation of small structures is central to modern science and technology. The most evident examples are in microelectronics, where smaller has meant less expensive, faster, more components per chip, higher performance and lower power needs.21,22 Micro-fabrication techniques, which have been developed for application in the electronics industry and in information technology, have been progressively employed in biomaterials research to study cell-surface interactions.23 One aim of these studies is the spatial control of cell attachment and organisation.24,25 The ability to generate patterns of proteins and cells on surfaces is important for biosensor technology,26 tissue engineering,27 and fundamental studies of cell biology,28 and can generally be called the field of nanobiotechnology. The positioning of biological ligands at well-defined locations

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on substrates is required for biological assays,29 for combinatorial screening,30 and for the fabrication of biosensors.31 Tissue engineering requires cells that are placed in specific locations to create organised structures in three dimensions. Furthermore, the growing demand in biosensor technology for high-density, high-sensitivity, multi-analyte chips can be met only with precise and reproducible patterning technologies that allow controlled positioning of chemically different active areas. The similarity of needs and limits between the implant and biosensor fields has led to the development of chemical, topochemical and topographical patterning methodologies that are applicable to both areas. Strategies that have been explored for fabricating patterned nanostructures take advantage of at least one of the following principles: · interaction of matter (`lithography') with photons (x-ray,32 ultraviolet33), energetic particles (electrons, ions, neutral metastable atoms) and scanning probes34 · replication against masters via physical contact35 · self-assembly of molecules36 and nanoscale objects.37 Most of these methods involve sophisticated instrumentation or treatments, and only a few of them are suited to create topographical variations at polymer surfaces without changing the chemistry. Photolithography is the most widely applied technology in the fabrication of microelectronic structures but this technique is expensive and too slow for the production of nanometre features. Self-assembly techniques provide the opportunity to produce features with a higher resolution down to a few nanometres. Micro-contact printing allows fast and reproducible production of chemical micro- and nanostructures with a resolution as good as 50 nm. Smaller dimensions are limited by the contact stability of the stamp with the substrate. Dip-pen nanolithography, blockcopolymer lithography, colloidal lithography, and self-assembled monolayer lithography are additional techniques that use self-assembly strategies to produce nanoscale structures. These methods are described in more detail below.

6.2

Lithographic patterning with photons, particles and scanning probes

Photolithography (Fig. 6.1(a)) and electron-beam (e-beam) lithography (Fig. 6.1(b)) are the most frequently used techniques in micro- and nanotechnology, used to manufacture components for computer technology. However, they have also been used most extensively to create features for patterning proteins and cells. Photolithography uses electromagnetic radiation to introduce a latent image into the appropriate material.38,39 The surface is first covered with a radiation sensitive film, usually a polymer called `resist', then exposed to a beam of radiation, which modifies polymer properties at the irradiated areas. Subsequently, a patterned structure can be developed through etching or

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6.1 (a) Schematic of the general principle of photolithography and patterning with self-assembled monolayers (SAMs) (e.g. on Au). A photoresist is deposited on the substrate (e.g. by spin-coating) of choice and light is shone through the mask of defined pattern creating a pattern. A SAM with one type of head group (thiol 1) is chemisorbed onto the exposed Au substrate. The leftover photoresist is then dissolved away and a second SAM (with a different head group) is chemisorbed into the gaps, thus creating a surface with different chemical patterns. (b) Schematic of lithography using either electrons (ebeam) or photons. The principle is the same as (a) except the pattern is written directly into the photoresist followed by developing with a solvent. In e-beam lithography features down to 5±10 nm are easily and reproducibly fabricated.

dissolution of the image, leaving a pattern of polymer on the surface that serves as a mask for further surface treatment of the uncoated areas. Different variations of these techniques, such as the use of short wavelength light sources, for example deep UV or X-ray, the chemical adjustment of the polymer resist material to the light source has allowed further decrease of patterned dimensions. The major drawbacks with conventional lithography are the large costs due to cumbersome fabrication processes and low throughput rates for nanometre sized features. In addition, photolithography is also not well suited for the modification of surfaces involving delicate biomolecules.

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6.2.1 Photolithography (smallest feature size obtainable is 50 nm) Photolithographic micro-fabrication of surfaces is a well-established technique for the production of model surfaces with defined topographies or with chemically defined patterns. Photolithography with photons uses a mask or a hole to localise the beam spatially. This technique is one of the most widely used in science and technology. The resolution of photolithography increases as the wavelength of the light used for exposure decreases. On the other hand, as structures become increasingly small, they also become increasingly difficult and expensive to produce. With 193 nm light from an ArF excimer laser, patterning of features down to 150 nm has been achieved.40 Patterning of features below 100 nm becomes challenging because of the lack of transparent materials suitable for lenses or for support for photomasks at wavelengths below 193 nm. Sun et al. developed a scanning near-field photolithography (SNP) technique to pattern self-assembled monolayers (SAMs) on gold with feature sizes of 40±50 nm (Fig. 6.2). These materials have been used as masks for pattern transfer to the underlying gold substrate by wet chemical etching creating three-dimensional architectures.41 Chemical modification of the terminal reactive group of patterned SAMs will enable development of materials of interest and find applications in diverse areas of nanotechnology.42 Photolithography and focused laser methods have been used to pattern surfaces with molecular layers,43,44 but photolithography requires the use of harsh solvents and bases, making it incompatible with many biological molecules. The laser method uses an interference technique that does not allow generation of patterns of arbitrary complexity. However, protein patterning using chemical linkers to create a heterogeneous monolayer is achievable.44 Silanes have been a widely used reagent to attach proteins to silica or metal surfaces because they can withstand the harsh solvent system required to remove the resist. Textor et al. describes a photolithographic technique based on evaporated metal films of titanium, aluminium, niobium and vanadium with patterns in the 50±150 lm size range to study protein and osteoblast cell interactions with different geometries and chemical composition of the patterns.45,46 Spatz et al. used ArF excimer laser ablation to generate selforganised, grating-like periodic structures on the surface of PET films.47 They demonstrated that both the morphology and the orientation of the melanocytes are determined by the topography of the PET substrates structured by the laser. A novel PEG surface modification and patterning process has been proposed.48 This approach mimics traditional photolithography in that spin coating and UV exposure through a photomask are employed to create a polymer pattern. The obtained PEG hydrogel microstructures resisted protein adsorption and cell adhesion whereas non-modified substrate surface promoted cell adhesion and spreading.49

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6.2 Generating nanostructures by scanning near-field photolithography (SNP). In this example a combination of self-assembled monolayers and wet chemical etching are used to create the patterns. (a) An AFM topography image reveals a set of 55  5 nm parallel trenches etched into a gold film. The SAM used was a HSC15CH3 molecule chemisorbed to the gold surface. The wet chemical etching was performed using a solution of Fe(CN)62+/Fe(CN)63+. (b) The cross-sectional topography trace orthogonal to the lines (taken from ref. 41, with permission).

Other problems accompanied with photolithography are damage and deformation of the mask on exposure to energetic radiation and the high cost and low speed of this technique. Furthermore, it cannot be easily applied to nonplanar surfaces, tolerates little variation in the materials that can be used and provides almost no control over the chemistry of the patterned surface. Plasma lithography Goessl et al. have developed a method for the patterned immobilisation of cell surface receptor ligands on biomaterials surface based on the combination of radio-frequency glow discharge plasma (RFGD) and photo-microlithography.50 The method called plasma lithography combines the high spatial resolution of photolithography and the chemical versatility of plasma surface modification. In the first step a non-fouling background is deposited onto the polymer surface, and the surface is then coated with a positive photoresist that is exposed to UV

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through a photomask that defines the pattern. In the second step, plasma treatment, a fluorocarbon plasma polymer is deposited onto this patterned surface. Surfaces composed of micro-patterned domains of fluorocarbon polymer can control the shape and size of vascular smooth muscle cells (SMC). Spreading restricted cells formed a well-ordered actin skeleton, and the cells were still confined to the areas of the adhesive pattern after two weeks. Dai et al. generated high-resolution surface patterns of various surface functionalities through the H2O-plasma etching using a TEM grid as a mask.51 They also extended this technique to include conducting polymers using the plasma-patterned, metal-sputtered substrates as the electrodes for electropolymerisation. The regions of the electrode surface covered by the plasma polymer are electrically insulating and hence inactive toward electropolymerisation, whereas the uncovered areas can effectively initiate electropolymerisation. Thissen et al. described a controlled excimer laser ablation technique of a plasma polymer-PEG modified substrate to fabricate surface patterns suitable for spatial control of protein adsorption and subsequent cell attachment.52 By using appropriate masks in the laser beam, a resolution of approximately 1 m was obtained.

6.2.2 Electron beam lithography (e-beam) (feature size obtainable down to 10 nm) In e-beam lithography, a focused electron beam is used to form patterned nanostructures in an electron sensitive resist film.53 Interaction of the electron beam with the resist causes changes in solubility and the formation of local spots that become soluble in a developer. Pattern formation using electrons is a method capable of forming patterns with nanometre resolution in a resist film. With e-beam writing resolution of 10 nm was achieved more then 20 years ago,54 conventional lithography with focused beams of electrons is slow and expensive.55 Consequently, e-beam lithography is more suitable for producing photomasks for optical lithography. Plastic surfaces have been patterned by combining homogeneous polymer grafting with e-beam irradiation and localised laser ablation of the grafted polymer,56 thus fibronectin was adsorbed selectively onto ablated domains and hepatocytes adhered specifically onto the ablated domains adsorbed with fibronectin.

6.2.3 Scanning probe lithography (SPL) (feature size obtainable down to 1 nm) Scanning probe lithography uses a small tip scanned near the surface of a sample to image and modify surfaces with atomic resolution.57,58,59 Advantages of SPL are the ability to generate features with any geometry as well as patterning of

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non-planar surfaces.60 Additional advantages of such an approach include simplicity (no need for clean room fabrication environment), the ability to pattern proteins under non-denaturing solution environments, and direct interrogation of protein binding events. Recently, SPL methods have been explored for patterning protein surfaces. Examples include the mechanical scraping of open patches within protein-resistant polymer films for subsequent protein adsorption,61 application of scanning electrochemical microscopy to derivatise electrode surfaces,62 and the use of nano-grafting to incorporate reactive sites into self-assembled monolayers.63 This method offers the possibility of creating as well as of reading extremely high-density protein arrays.64 SPL is used to scratch nanostructure in soft materials,65 to change the head group or packing density of organic monolayer catalytically66 or to write 30 nm patterns of alkanethiols on gold.67 SPL is a slow method because the writing speed is limited by the mechanical resonance of the tip and piezoelectric elements that maintain the constant separation between a tip and sample. For that reason SPL is better suited for the formation of masters than for replication, but efforts to develop faster SPL are continuing. The use of multiple-tip systems has increased production speeds remarkably.68

6.2.4 Dip-pen nanolithography (DPN) (feature size obtainable less than 100 nm) In dip-pen nanolithography (DPN), an atomic force microscope (AFM) tip is inked with a material known to self-assemble on a solid substrate, as depicted in Fig. 6.3. The tip is brought into contact with the surface, a water meniscus is formed and adsorbed ink is transferred to the substrate when the probe is held in contact or moved along the surface.69,70 The first DNP investigations used thiolbased ink and gold substrates67 but later it has been extended to other inks and substrates. Various oligonucleotides functionalised with a hexanethiol linker have been inked on gold.71 Collagen was also written on a gold surface by cysteine binding on the gold without destroying the helical structure of the collagen.72 Thiolated collagen and collagen-like peptide has been patterned on gold with line widths as small as 30±50 nm, the largest molecule thus far positively printed on a surface at such small length scales. Moreover, DPN can be used to generate complex multi-component nano-arrays of native proteins that are biologically active and capable of recognising a biological complement in solution.73,74 Other inks that do not chemisorb to the surface are being developed. Fluoroscently labelled proteins can be written on modified glass surfaces.75 Utilising electrostatic interaction between positively charged parts of a protein and negatively charged silica, proteins have been directly written onto silica surfaces.76,77 Hyun et al. developed a method to fabricate chemically reactive

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6.3 The concept of dip-pen nanolithography (DPN). Schematic of the process using an AFM tip to deposit chemical moieties on surfaces after it is wetted with the desired molecule of choice. (Taken from Ref 67, with permission).

nanoscale features by patterning a SAM of a COOH-terminated alkanethiol on a gold substrate by DPN, followed by covalent immobilisation of biotin on the nano-patterned SAM and subsequent molecular recognition of streptavidin from solution.78 The resulting streptavidin nano-pattern provides a universal platform for molecular recognition-mediated protein immobilisation because of the ubiquity of biotin-tagged molecules. By this method, periodic arrays of biotinBSA with feature size of 230 nm were readily fabricated.

6.3

Soft lithographic techniques

Current lithography techniques appear to have reached their limits in terms of minimal feature size and fabrication cost. Thus, alternative methods need to be explored to go beyond these limits. An attractive candidate for achieving nanoscale structures is the phenomenon of self-organisation, in which certain materials under certain conditions will arrange themselves or self-assemble into stable, well-defined structures, e.g., when deposited onto a surface.79,80 The development of self-assembled systems has been a major advance in material fabrication technology during the last ten years.81,82 Self-assembly concepts originate from biological processes such as the folding of proteins, the formation of the DNA double helix and the formation of cell membranes from phospholipids. In self-assembly, subunits spontaneously organise into stable, well-defined structures based on non-covalent associations. Self-assembled monolayers

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(SAMs) provide well-defined ordered structures and chemistries that can be systematically varied. Also, spatially defined arrays of SAMs can be prepared by combining self-assembly with patterning methods such as micro-contact printing83 and photolithography.84 Self-assembly is being examined for patterning at scales greater than 1 m and applications are moving to smaller dimensions. Examples of self-assembly as a method for fabricating nanostructures include micro-contact printing of self-assembled monolayers,85 reactive ion etching with thin films of block copolymers as a mask86 and the synthesis of mesoporous materials with aggregates of surfactants as templates.87 Typically, this is a non-serial process (since the whole system organises itself at the same time) allowing patterning of large areas with high short-range and some long-range order. The thickness of the SAM is usually 2±3 nm, and can be tuned with a accuracy of 0.1 nm by varying the number of carbon atoms in say the monolayer alkyl chain. Polymers and block copolymers are an interesting class of molecules for nanomaterials fabrication88 and show a broad variety of self-organising patterns; such polymeric nanostructures find applications in diverse areas that include optics, biochemistry, and material science. The ability to use different molecules with well-defined chemical end groups adds to the flexibility of this approach to create nanoscale patterns of specific functionalities. Soft lithography has been developed as an alternative to photolithography and comprises a number of lithographic techniques such as micro-contact printing (CP),89 replica moulding (REM),90 micro-transfer moulding (TM),91,92 micromoulding in capillaries (MIMIC),93 and solvent-assisted micro-moulding (SAMIM).94 In soft lithography, an elastomer is cast against a rigid master (silicon wafer) and the elastomeric replica subsequently used as the stamp, giving structures that can be as small as 100 nm. Masters are made using conventional high-resolution nano-lithographic techniques. In this way it is possible to produce multiple copies of indistinguishable nanostructures from a single master, rapidly and economically.95 The common feature of these techniques is the use of a patterned elastomer (usually poly(dimethylsiloxane (PMDS)) as the mould, stamp or mask to generate or transfer a pattern to the substrate. In addition, soft lithography uses flexible organic molecules and materials rather than rigid inorganic materials commonly used in photolithography. Table 6.1 compares the features of photolithographic and soft lithographic approaches to surface patterning.83 In CP a PMDS stamp is used to transfer molecules of the `ink' to the surface of the substrate by contact, whereas in REM a PMDS polymer is used to duplicate structures of a master into multiple copies. In TM a drop of liquid pre-polymer is applied to the patterned surface of a PMDS mould. The filled mould is placed in contact with a substrate and irradiated or heated. After curing of the liquid precursor, the mould is peeled away to leave a patterned microstructure on the substrate surface. In MIMIC a PMDS mould in conformal contact with a substrate surface forms a network of empty channels

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Table 6.1 Comparison of photolithography and soft lithography approaches83 Photolithography

Soft lithography

Definition of patterns

Rigid photomask (patterned Cr)

Elastomeric stamp or mould (PDMS)

Materials that can be patterned directly

Photoresist

Photoresist

SAMs

SAMs Polymers Colloidal materials Sol-gel materials Organic and inorganic salt Biological macromolecules ca. 30 nm

The smallest feature size

ca. 50 nm

that can be filled with a pre-polymer. After curing the liquid into a solid, the PMDS mould is removed leaving a network of polymeric material on the surface of the substrate. SAMIM is a combination of replica moulding and embossing. A PMDS mould is wetted with a solvent, and is brought into contact with the surface of the polymer. The solvent dissolves a thin layer of the polymer, and the resulting gel-like fluid conforms to the surface topology of the mould. After evaporation of the solvent a relief structure is formed on the polymer surface. Soft lithographic techniques require little capital investment and are very simple. They can often be carried out under ambient laboratory conditions. They are not subject to the limitations determined by optical diffraction and optical transparency. Table 6.2 is a summary of soft lithographic and other nonphotolithographic approaches to creating patterned substrates.91,94,96-110 Producing patterned biomolecule layers using soft lithographic techniques has a specific importance to a number of growing techniques such as advanced tissue engineering, biomineralisation, DNA computing and cultured neural networks. It is promising for microfabrication of relatively simple, single-layer structures for use in cell culture, in sensors or as microanalytical systems.111,112

6.3.1 Microcontact printing (CP) (feature size obtainable down to 35 nm) CP is also a method for chemically and molecularly patterning surfaces down to sub-micrometre length scales. It has been used to pattern self-assembled monolayers (SAMs) of compounds such as hexadecanethiols and octadecytrichlorosilane (OTS) on gold113 and SiO2114,115 surfaces. CP has an advantage over conventional photolithographic patterning methods in that it requires no harsh chemicals making it suitable for patterning biologically active layers.

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Table 6.2 Summary of non-photolithographic methods83 Method

Resolution

Reference

Injection moulding Embossing (imprinting) Cast moulding Laser ablation Laser-induced deposition Electrochemical micromachining Ink-jet printing Stereolithography Soft lithography Microcontact printing ( CP) Replica moulding (REM) Microtransfer moulding ( TM) Micromoulding in capillaries (MIMIC) Solvent-assisted micromoulding (SAMIM)

10 nm 25 nm 50 nm 70 nm 1 m 1 m 50 m 50 m

96, 97 98, 99 100 101 102 103 104 105 106, 107 108 109 91 110 94

35 nm 30 nm 1 m 1 m 60 nm

Additionally, it should be possible to pattern multiple molecular layers by repeated application of CP. CP was mainly developed with SAMs of alkanothiol on gold113 and silver.116 The CP technique is very simple; a PMDS stamp is first inked with a solution of molecules, often proteins or alkanethiols, and brought into contact with the surface of a substrate. The soft PDMS stamp makes conformal contact with the surface, and molecules are transferred directly from the stamp to the surface in the time frame of a few seconds. After printing, unmodified regions can be back-filled by deposition of a second molecule. Figure 6.4 is a schematic representation of the CP process. A key requirement for successful CP is the appropriate surface chemistry. The surface chemistry should allow high spatial resolution, low background adsorption, and high selectivity for molecule immobilisation. These conditions can be accomplished by using SAMs. If suitable SAMs are chosen, marked contrast in properties such as wetting or protein binding may be obtained. For example, Prime et al. used CP to generate a pattern of hydrophobic alkanethiols sites where proteins deposited from solution while non-patterned regions were blocked with PEG to prevent protein adsorption.117 They extended this method by combining CP and chemical reaction of ligands containing both amino and anhydride groups providing patterned SAMs presenting several ligands on the same surface, especially polar, charged, or structurally complex groups such as proteins, polymers, or oligosaccharides.118,119 Patterned SAMs on gold can also be used as ultra-thin resists in selective wet etching, protecting coated areas from chemical etching, while uncoated gold areas are removed to leave a structured gold film.120 The ability to pattern SAMs by CP, and resulting control over the adsorption of adhesive proteins, enables the patterning of cells on substrates121 and enables change in the size and shape of cells.122

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6.4 Schematic of the well-established microcontact printing process developed by Prof. Georges Whitesides' group at Harvard University (taken from ref. 91, with permission).

However, this strategy cannot easily be extended to other non-metallic surfaces such as silicon or polymers. The importance of using patterning of silicon surfaces is immense for microelectronics, micro-machinery and micro-fluidics. Zhu et al. developed a two-step alkoxy monolayer assembly method for CP used directly on silicon surfaces based on a reaction between alcohol functional groups and Cl-terminated silicon surfaces.123 In an extended method they produced patterns of HO- and CH3O-terminated PEG regions whereby the HOterminated PEG monolayer was activated to immobilise protein molecules covalently.124 Generally, CP is inexpensive, very rapid and a powerful method for surface structuring, also applicable to curved substrates or inner surfaces and able to be used over a relatively large surface area (50 cm2) in a single step.125 Repeated printing using different stamps can allow complex surface patterns to be made of more than one kind of molecule. The smallest features generated to date with CP are trenches etched in gold with lateral dimensions of approximately 35 nm.126 By manipulation of the elastomeric stamps and the chemistry of SAMs formation, it is conceivably possible to reduce the size of features generated even further.127 Patterned microstructures fabricated using this technique may be useful in microelectronics, for micro-analytical systems, sensors, solar cells, and optical components. Patterned SAMs have also been used for the control of extracellular matrix protein adsorption and for the attachment of cells.128 It is possible to control the shape of the cell that attaches to a surface by creating islands of cell

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adhesion proteins onto substrate surface and thus to control cell growth and morphology. Polymers are widely used as biomaterials but cannot be modified by SAMs. Therefore, Chilkoti et al. developed a new method called microstamping on an activated polymer surface (MAPS) that enables patterning of biological ligands and proteins on polymer surfaces with a spatial resolution of at least 5 m.129,130 A requirement for MAPS is the presence of reactive functional groups in the polymer. Therefore, the surface of a polymer is modified to introduce the reactive group of interest. In a second step, an elastomeric stamp is inked with a biomolecule containing a complementary terminal reactive group and brought into contact with the activated polymer surface. This results in spatially resolved transfer and covalent coupling of the biomolecule on the polymer surface. To demonstrate the concept of MAPS, they patterned an amine terminated biotin molecule on carboxylated PET.131 Because MAPS is a simple and flexible method, it will be applicable to a wide variety of polymers that are amenable to surface modification. Hyun et al. demonstrated a simple method of patterning cells on commonly used polymeric biomaterials in the presence of cell adhesion proteins.132 This approach involves coating a polymer surface by solution casting a chemically reactive, biologically non-fouling comb polymer presenting short oligoethylene glycol side chains, activation of the surface, and patterning a cell adhesive peptide onto the surface of the comb polymer by MAPS to spatially direct the interactions of anchorage dependent cells with the substrate. The patterns of fibroblasts and endothelial cells were stable for long periods of time depending on the spacing distance between isolated features such as a square or stripe.133 Csucs et al. introduced a new variant of CP based on adsorption of a polycationic graft copolymer, poly-L-lysine-g-poly(ethylene glycol) (PLL-gPEG), on negatively charged surfaces through electrostatic interactions, rendering them highly protein and cell resistant.134 Furthermore, PLL-g-PEG can be modified by introducing biologically active ligands such as short cell adhesion peptide sequences (viz. RGD) at the terminus of PEG chains. Printing cell adhesive, peptide modified PLL-g-PEG followed with backfill with cell repulsive non-functionalised PLL-g-PEG, allows control of the concentration of biomolecules presented to cells at the interface. If the separation distance between adhesion sites was large (at least 30 m spacing) fibroblasts were found primarily on printed, cell-adhesive regions, whereas with smaller spacing between adhesive sites (less then 30 m) the cells were able to occupy nonadhesive regions as well. This observation can be explained by a bridging effect where cell pairs or agglomerates grow over non-adhesive regions without formation of cell-surface contact. Nanoscale objects such as colloidal particles can also be patterned on a variety of substrates including glass, silicon, and polymers using CP.135 An important feature of CP is that many proteins retain their biological activity after printing.136 It is possible to print the same surface with several

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proteins using different stamps having various inks printed many times or using parallel inking of a stamp followed by a single printing.137 However, two limitations of CP as presently used hinder its progress in printing arbitrary patterns of biological molecules. First, the PDMS elastomer used for the stamp in CP is hydrophobic. When printing with water-based, biological solutions, the poor wetting characteristic of the hydrophobic stamp yields extremely non-uniform patterns. Therefore, the hydrophilisation of PDMS is a requirement for employing biomolecules and polar inks for CP that do not have an affinity for native PDMS. The simplest method for preparing a hydrophilic stamp is oxidation of the PDMS surface by O 2 -plasma treatment,138,139 but the stamp does not remain hydrophilic for a long time because of the migration of low-molecular silicone residues from the bulk to the air-stamp interface.140 Delamarche et al. overcame this problem by grafting of a PEG silane onto plasma oxidised PDMS that resulted in a stable hydrophilic stamp surface for an extended period of time.141 The second limitation is that the soft PDMS stamp lacks the rigidity necessary for precision alignment and geometrical control of the pattern and due to stamp deformation is not suited for the printing of features below 250 nm. To resolve this problem, James et al. developed a thin elastomeric stamp on a rigid glass backing in order to allow printing of small isolated features and improve the alignment and geometrical control of the patterns.142 Another, often neglected, drawback of PDMS-based CP is the low molecular weight PDMS contamination presented on stamped surfaces.143 To overcome mechanical and contamination problems, Csus et al. investigated the possibility of new materials in CP applications. A copolymer of ethylene and a -olefin such as butene or octene (POPs) was used for printing protein patterns.144 Due to the higher bulk modulus of the polyolefin stamp, much higher printing quality in submicrometre ranges can be achieved with the POPs stamp than with the PMDS stamp. Also, multiple protein patterning is difficult to achieve; a method for printing different proteins onto a single substrate involves using stamps having different patterns or various inks that are printed many times onto the same substrate. It has been found that accurate reproduction of patterns realised in PDMS stamps on gold substrates is problematic on a scale smaller than 500 nm due to the diffusion of ink molecules from contacted to non-contacted areas.145 Li et al. used CP without ink, and obtained sub-micrometre edge resolution (less than 100 nm).146 They brought a surface oxidised PDMS stamp into contact with SAMs of acid labile adsorbates. The silicon oxide layer on the outer surface of the stamp was acidic enough to hydrolyse these adsorbates. Finally, scanning probe contact printing is a method for generating patterns less than 500 nm using a scanning probe with an integrated elastomeric tip to transfer chemical materials onto a substrate.147 Each contact printing action creates an individual dot. This method allows generation of structures made of organic monolayers on surfaces with excellent multi-ink and alignment registration capabilities.

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6.3.2 Micro-fluidic techniques (micro-moulding in capillaries ± MIMIC) In CP the deposition of liquid samples followed by drying can lead to poorly defined patterns, protein aggregation and the loss of biological activity. An alternative method to bring a controlled amount of biomolecule in contact with a surface without loss of biological activity is through use of a micro-fluidic network (-FN). The method is based on conformal contact ± a watertight seal that forms via van der Waals forces ± between a soft PDMS structure and a hard surface such as silicon, glass or some polymer. One or both of the surfaces is structured so that a network of fluidics channels is formed when the two surfaces are brought together. The channel network can guide solutions of proteins or other molecules over the substrate and the molecules adsorb onto the surface of the fluidics channels with the pattern defined by the micro-fluidic network (Fig. 6.5a).148 Once adsorption is complete, the PDMS is peeled off, to give a patterned surface with molecules deposited on surface regions exposed to the solution. Microfluidic networks are a very powerful tool for surface structuring. This method is well suited for patterning highly sensitive entities like proteins and cells on a variety of substrates, and can be used at room temperature with aqueous solutions. It is a simple technique that forms patterned structures in a single step with very high accuracy on a wide variety of surfaces ± flat or curved ± with structures as small as 100 nm. Different molecules can be made to flow in different channels of the network. Delamarche et al.110 used MIMIC to pattern a variety of substrates with immunoglobulins with submicron resolution (Fig. 6.6(a)Ð(c)). Only microlitres of reagent were required to cover several mm2 sized areas. This technique enabled simultaneous and highly localised immunoassay for the detection of different immunoglobulins. The independent network of capillaries allows simultaneous attachment of different biomolecules in each zone of flow, as shown in Fig. 6.6(b).149 Gradients can also be made using solutions at low concentrations. The high surface-to-volume ratio of the small channels results in a depletion of the solution as it travels through the network, resulting in lower and lower quantitites of molecules deposited on the channels walls. Patel et al. developed a micro-fluidic patterning technique for any biotinylated ligand to be patterned onto the surface of a biodegradable polymer. Spatial control over cell appearance was also observed when using these templates to culture endothelial and nerve cells. Furthermore, pattern features containing laminin peptide sequences were achieved for directional control of nerve regeneration.150 However, MIMIC cannot form isolated structures or patterns on contoured surfaces because it forms a hydraulically connected network of capillaries. Additionally, the extremely slow filling of small capillaries may limit the usefulness of MIMIC in certain types of nanofabrication

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6.5 The process of -FN for patterned delivery of proteins. (a) Patterned elastomer that forms a FN by contact with a substrate allows the local delivery of a solution of biomolecules to the substrate. (b) Flow of liquid between the filling pad and an opposite pad fills the array of microchannels that constitute the strategic part of this device. (c) Assembly of different zones of flow on the surface results from the independence of capillaries, each requiring only a small volume (1 l) of liquid to fill the zone and derivatise the underlying substrate. Left panel, top view; right panel, side cut along the channel (taken from ref. 110, with permission).

applications. Micro-fluidic networks must be sufficiently hydrophilic to promote filling of micro-channels by capillary action and should incorporate proteinrepellent surfaces to prevent depletion of proteins by adsorption to the walls of micro-channels. One approach to making micro-fluidic networks wettable and protein-repellent is to functionalise the channels with PEG.151 Another approach has been to prepare a polymer-based micro-fluidic device with defined and chemically reactive interfaces by deposition of sub-micrometre thin reactive coatings on the interior surface of micro-fluidic devices. The resulting coating was then used for immobilisation and self-assembly of a variety of biological ligands, proteins, and cells.152

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6.6 Schemes for the delivery and attachment of two different IgGs using a FN (a) followed by an immunoassay for the attached proteins after removal of the FN (b). A composite digital image shows light emitted from fluorescently tagged antispecies IgGs, each specifically recognising its binding partner previously patterned on a glass surface (c). The immunoassay is carried out with a heterogeneous solution of IgGs: tetramethyl rhodamine isothiocyanateconjugated antibody to chicken IgG (red), fluorescein isothiocyanateconjugated antibody to mouse IgG (green), and R-phycoerythrin-conjugated antibody to goat IgG (orange-red), each diluted 1:300 from their concentrated solutions. The left stripe comprises chicken IgGs and the right stripe comprises mouse IgGs. No light was evident from nonspecific deposition of antibodies to goat IgG anywhere on the surface. There was no green fluorescence on the left channel or red fluorescence on the right channel; each colour channel was collected independently so that such emission, resulting from cross-reactivity between the antibodies or their uncontrolled deposition, would have been easy to detect.

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6.4

Colloidal-based fabrication techniques

Colloidal particles as lithographic masks can increase the speed of patterning surfaces with nano-sized features. Colloidal particles of different materials can be produced with mono-disperse size distribution down to nanometre length scales. The colloidal particles served as an etch mask for the underlying substrate leaving the areas under the particles as hills. These particles can also be used as lift-off masks by filling up the surface surrounding the particles with a thin film. After removal of the particles, the underlying surface is exposed where the particles are located. One of the main advantages of this method over other lithographic techniques is that complex sample shapes can be patterned or coated with nano-porous layers. Spatz et al. have developed a copolymer micelle nanolithography method for generation of nanostructures on conductive and isolating substrates. The pattern dimension and geometry is controlled by the combination of the self-assembly of block copolymer micelles with pre-structures formed by photo- or e-beam lithography.153,154 This method allows bridging the length scale between several nm and 200 nm, through linking self-assembly nanostructures to structure sizes available from photo- or e-beam lithography. E-beam and photolithography by themselves are not able to write structures with such small dimensions over large areas, whereas self-assembly methods, such as the formation of nanoparticles in block polymer micelles cannot position particles in patterns with large ( m) separation distance and aperiodic arrangements. After formation of photo or ebeam resist patterning and coating with micelles loaded with a metal precursor, the substrate is dipped in a solvent to remove the resist film and any micelle deposited on top of the resist. The micelles inside the holes or grooves remain adsorbed at exactly the position where the hole originally formed. The substrate is treated with oxygen plasma in order to remove the polymer selectively, leaving behind Au-nanoparticles loaded into the core resembling the lithographic pattern. Combining these two techniques, gold dots of 1 nm diameter can be positioned with high accuracy. These dots can act as an anchor for individual macromolecules and biomolecules. The Au particles have been used to bind streptavidin proteins in an ordered array.155 Furthermore, a wafer with receptor modified gold dots may be used for the separation, location and screening of DNA or proteins. In order to group a small number of Au-nanoparticles, locally monomicellar layers were exposed directly to a focused electron beam so that a micelle located in an exposed area was chemically modified and immobilised on the substrate surface. This indicates that monomicellar layers can be applied as a negative ebeam resist. Au-nanoclusters are then deposited from diblock copolymer micelles by hydrogen plasma treatment. By this technique `micrometre nanostructured' patterns can be created marked by uniform 2, 5, 6, or 8 nm Au clusters.156

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Finally, colloidal lithography based on deposition of polystyrene particles onto flat oxidised titanium surfaces for particle sizes between 20 and 500 nm has been demonstrated. Methods to minimise aggregation with large particles or with high surface coverage has involved heating; co-deposition of silica particles and heating were shown to optimise the patterned protocols.157 Frey et al. developed a method termed ultra-flat nanosphere lithography (UNSL) that combines nanosphere lithography and ultra-flat template stripping to create 150 nm features of Au, Al, or Ag embedded in a matrix of Al, Au, or SiOx with topographical variation of less than 1 nm between patterned features and matrix.158 In UNSL, a material is deposited onto mica through a mask created by a close-packed monolayer of nanospheres. After removal of the spheres, a second material is deposited onto the nanostructures. Subsequently, upon removing the mica, the surface in contact with mica reveals flat nanostructures of the first material embedded in a matrix of the second deposited material. This technique enables the fabrication of chemically distinct patterns of one material embedded in a matrix of another material with minimal topographical variation. Independent control of chemistry and topography is important for studies of cell-surface interactions.

6.5

Template-imprinted nanostructured surfaces

The technique of molecular imprinting creates specific recognition sites in polymers by using template molecules.159,160 Molecular recognition is attributed to binding sites that complement molecules in size, shape and chemical functionality. Attempts to imprint proteins have been made with limited success.161,162 Shi et al. report a multi-step method for imprinting surfaces with protein-recognition sites using radio-frequency glow discharge plasma deposition to form polymeric thin films around proteins that were coated with disaccharide molecules. The disaccharides become covalently attached to the polymer film, creating polysaccharide-like cavities that exhibit highly selective recognition for a variety of template proteins.163 Figure 6.7 shows the fabrication process to create nanostructures or imprints that selectively adsorb individual proteins from a complex mixture.

6.6

Conclusion

As in physics and electronics, the multidisciplinary field of nanobiotechnology, which includes materials science, chemistry, biology and medicine, is becoming more reliant on the fabrication of structures and patterns with dimensions below 100 nm. To be useful in an industrial setting, ideally methods for nano-patterning and fabrication methods will need to be of low cost, flexible in terms of materials used, high precision during and after replication, be able to pattern on non-planar substrates, be produced with high

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6.7 Protocol for template imprinting of proteins. (a) Template protein was adsorbed onto a freshly cleaved mica in citrated phosphate-buffered saline (CPBS), pH 7.4. A 1±10 mM solution of disaccharide was spin-cast to form a 10±50 Ð sugar overlayer. The sample was put into the in-glow region of a 13.56 MHz RFGD reactor. Plasma deposition of C3F6 was conducted at 150 mtorr and 20 W for 3±6 min, forming a 10±30 nm fluoropolymer thin film. The resulting plasma film was fixed to a glass coverslip using epoxy resin and oven cured. Mica was peeled off and the sample was soaked in a NaOH/NaClO (0.5/1.0%) solution for 0.5±2 h for dissolution and extraction of protein. A nanopit with a shape complementary to the protein was created on the imprint surface. (b) A tapping mode AFM image of the surface of a fibrinogen imprint, together with a drawing of fibrinogen. (c) Mechanisms for the specific protein recognition of template-imprinted surfaces. A nanocavity-bound template protein is prevented from exchange with other protein molecules in the solution because of steric hindrance and an overall strong interaction; the latter is due to many cooperative weak interactions, involving hydrogen bonds, van der Waals forces and hydrophobic interactions for example (taken from ref. 163, with permission).

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speed and throughput and allow potentially for 3-D fabrication.164 Because all of these requirements may never be met by a single fabrication method, it may be necessary to create hybrid techniques which combine several of the methods described above. In particular, the present patterning techniques have proved to be well suited for specific applications, but do suffer from a number of limitations. Current photolithographic techniques that do not use a stamp require complex chemistry, while solvents used in conventional photolithography may denature or degrade the deposited layer. Lithographic techniques are costly because of the elaborate fabrication processes as well as the low manufacturing throughput. While nonlithographic techniques allow spatial control, the use of elastomeric stamps hinder reproducibility over large areas, transfer contaminants and degrade over time. Molecular self-assembly allows us to push the limits of soft lithography to the molecular scale. Similarly, creating boundary conditions by lithographic approaches allows direct self-assembly in more promising directions. The combination of both approaches, both powerful at their respective length scales, has large potential for future applications. In summary, there is great potential in the combination of self-assembly and conventional lithographic techniques. Self-assembly guarantees that the smallest sized features and functionalities are programmed into the structure and chemistry of molecules, while conventional lithography meets the demands of controlled structure location of nanostructures with large separation length and also aperiodic arrangements.

6.7

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Whitesides G M, Rich A, `Biological surface engineering: a simple system for cell pattern formation' Biomaterials, 1999, 20, 1213±1220. Xia Y N, Whitesides G M, `Soft lithography' Annu. Rev. Mater. Sci., 1998, 28, 153±184. Wilbur J L, Kumar A, Biebuyck H A, Kim E, Whitesides G M, `Microcontact printing of self-assembled monolayers: Application in microfabrication' Nanotechnology, 1996, 7, 452±457. Whitesides G M, Xia Y N, `Replica molding: Complex optics at lower costs' Photon Spectra, 1997, 31, 90±91. Delamarche E, Bernard A, Schmid H, Michel B, Biebuyck H, `Patterned delivery of immunoglobulins to surfaces using microfluidic networks' Science, 1997, 276, 779±781. Brittain S, Paul K, Zhao X M, Whitesides G M, `Soft lithography and microfabrication' Phys. World, 1998, 11, 31±36. Qin D, Xia Y N, Rogers J A, Jackman R J, Zhao X M, Whitesides G M, `Microfabrication, microstructure and microsystems' Top Curr. Chem., 1998, 194, 1±20. Kumar A, Biebuyck H A, Whitesides G M, `Patterning self-assembled monolayersapplications in materials science' Langmuir, 1994, 10, 1498±1511. Xia Y N, Mrksich M, Kim E, Whitesides G M, `Microcontact printing of octadecylsiloxane on the surface of silicon dioxide and its application in microfabrication' J. Am. Chem. Soc., 1995, 117, 9576±9577. St. John P M, Craighead H G, `Microcontact printing and pattern transfer using trichlorosilanes on oxide substrates' Appl. Phys. Lett., 1996, 68, 1022±1024. Xia Y N, Kim E, Whitesides G M, `Microcontact printing of alkanethiols on silver and its application in microfabrication' J. Electrochem. Soc., 1996, 143, 1070± 1079. Prime K L, Whitesides G M, `Self-assembled organic monolayers- model systems for studying adsorption of proteins at surfaces' Science, 1991, 252, 1164±1167. Yan L, Zhao X M, Whitesides G M, `Patterning a preformed, reactive SAM using microcontact printing' J. Am. Chem. Soc., 1998, 120, 6179±6180. Lahiri J, Ostuni E, Whitesides G M, `Patterning ligands on reactive SAMs by microcontact printing' Langmuir, 1999, 15, 2055±2060. Geissler M, Schmid H, Bietsch A, Michel B, Delamarche E, `Defect-tolerant and directional wet-etch systems for using monolayers as resists' Langmuir, 2002, 18, 2374±2377. Mrksich M, Dike LE , Tien J, Ingber D E, Whitesides G M, `Using microcontact printing to pattern the attachment of mammalian cells to self-assembled monolayers of alkanethiolates on transparent films on gold and silver' Exp Cell Res, 1997, 235, 305±313. Singhvi R, Kumar A, Lopez G P, Stephanopoulos G N, Wang D, Whitesides G M, Ingber D E, `Engineering cell shape and function' Science, 1994, 264, 696±698. Jun Y, Le D, Zhu X Y, `Microcontact printing directly on the silicon surface' Langmuir, 2002, 16, 6766±6772. Jun Y, Cha T, Guo A, Zhu X Y, `Patterning protein molecules on poly(ethylene glycol) coated Si(111)' Biomaterials, 2004, 25, 3503±3509. Xia Y N, Venkateswaran N, Qin D, Tien J, Whitesides G M, `Use of electroless silver as the substrate in microcontact printing of alkanethiols and its application in

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microfabrication' Langmuir, 1998, 14, 363±371. 126. Biebuyck H A, Larsen N B, Delamarche E, Michel B, `Lithography beyond light: Microcontact printing with monolayer resist' IBM J. Res. Dev., 1997, 41, 159±170. 127. Xia Y N, Whitesides G M, `Extending microcontact printing as a microlithographic technique' Langmuir 1997, 13, 2059±2067. 128. Chen C S, Mrksich M, Huang S, Whitesides G M, Ingber D E, `Geometric control of cell life and death' Science, 1997, 276, 1425±1428. 129. Yang Z, Chilkoti A, `Microstamping of a biological ligand onto an activated polymer surface' Adv. Mater., 2000, 12, 413±417. 130. Hyun J, Zhu Y, Liebmann-Vinson A, Beebe T B, Chilkoti A, `Microstamping on an activated polymer surface: Patterning biotin and streptavidin onto common polymeric biomaterials' Langmuir, 2001, 17, 6358-6367. 131. Yang Z, Belu A M, Liebmann-Vinson A, Sugg H, Chilkoti A, `Molecular imaging of a micropatterned biological ligand on an activated polymer surface' Langmuir, 2000, 16, 7482±7492. 132. Hyun J, Ma H, Banerjee P, Cole J, Gonsalves K, Chilkoti A, `Micropatterns of a cell-adhesive peptide on an amphipilic comb polymer film' Langmuir, 2002, 18, 2975±2979. 133. Hyun J, Ma, H, Zhang Z, Beebe T P, Chilkoti A, `Universal route to cell micropatterning using an amphipilic comb polymer' Adv. Mater., 2003, 15, 576±579. 134. Csucs G, Michel R, Lussi J W, Textor M, Dansuer G, `Microcontact printing of novel co-polymers in combination with proteins for cell-biological applications' Biomaterials, 2003, 24, 1713±1720. 135. Hidber P C, Helbig W, Kim E, Whitesides G M, `Microcontact printing of palladium colloids: Micron-scale patterning by electroless deposition of copper' Langmuir, 1996, 12, 1375±1380. 136. Bernard A, Delamarche E, Schmid H, Michel B, Bosshard H R, Biebuyck H, `Printing patterns of proteins' Langmuir, 1998, 14, 2225. 137. Bernard A, Renault J P, Michel B, Bosshard H R, Delamarche E, `Microcontact printing of proteins' Adv. Mat., 2000, 12, 1067±1070. 138. Ferguson G S, Chaudhury M K, Biebuyck H A, Whitesides G M, `Monolayers of disordered substrates: Self-assembly of alkyltrichlorosilanes on surface-modified polyethylene and polydimethylsiloxane' Macromolecules, 1993, 26, 5870±5875. 139. Schmalenberg K E, Buettner H M, Uhrich K E, `Microcontact printing of proteins on oxygen plasma-activated poly(methyl mathacrylate)' Biomaterials, 2004, 25, 1851±1857. 140. Hillborg H, Gedde U W, `Hydrophobicity recovery of polydimethylsiloxane after exposure to corona discharges' Polymer, 1998, 39, 1991±1998. 141. Delamarche E, Donzel C, Kamonuah F S, Wolf H, Geissler M, Stutz R, Winkel PS, Michel B, Mathieu H J, Schaumburg K, `Microcontact printing using poly(dimethylsiloxane) stamps hydrophilized by poly(ethylene oxide) silanes' Langmuir, 2003, 19, 8749±8758. 142. James C D, Davis R C, Kam L, Craighead H G, Isaacson M, Turner J N, Shain W, `Patterned protein layers on solid substrates by thin stamp microcontact printing' Langmuir, 1998, 14, 741±744. 143. Yang Z P, Belu A M, Leibmann-Vinson A, Sugg H, Chilkoti A, `Molecular imaging of a micropatterned biological ligand on an activated polymer surface' Langmuir, 2000, 16, 7482±7492.

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144. Csus G, Kuenzler T, Feldman K, Robin F, Spencer N D, `Microcontact printing of macromolecules with submicrometre resolution by means of polyolefin stamps' Langmuir, 2003, 19, 6104±6109. 145. Delamarche E, Schmid H, Bietsch A, Larsen N B, Tothuizen H, Michel B, Biebuyck H, `Transport mechanisms of alkanethiols during microcontact printing on gold' J. Phys. Chem. B, 1998, 102, 3324±3334. 146. Li X M, Peter M, Huskens J, Reinhoudt D N, `Catalytic microcontact printing without ink' Nano Letters, 2003, 3, 1449±1453. 147. Wang X, Ryu K S, Bullen D A, Zou J, Zhang H, Mirkin C A, Liu C, `Scanning probe contact printing' Langmuir, 2003, 19, 8951±8955. 148. Delamarche E, Bernard A, Shmid H, Michel B, Biebuyck H, `Patterned delivery of immunoglobulins to surfaces using microfluidics network' Science, 1997, 276, 779±781. 149. Delamarche E, Bernard A, Bietsch A, Schmid H, Michel B, Biebuyck H A, `Microfluidic networks for chemical patterning of substrate: Design and application to bioassays' J. Am. Chem. Soc., 1998, 120, 500±508. 150. Patel N, Padera R, Sanders G H W, Cannizzaro S M, Davies M C, Langer R, Roberts C J, Tendler S J B, Williams P M, Shakesheff K M, `Spatially controlled cell engineering on biodegradable polymer surfaces' FASEB J., 1998, 12, 1447± 1454. 151. Papra A, Bernerd A, Juncker D, Larsen N B, Michel B, Delamarche M, `Microfluidic networks made of poly(dimethylsiloxane), Si, and Au coated with polyethylene glycol for patterning proteins onto surfaces' Langmuir, 2001, 17, 4090± 4095. 152. Lahann J, Balcells M, Lu H, Rodon T, Jensen K F, Langer R, `Reactive polymer coatings: a first step toward surface engineering of microfluidic devices' Anal. Chem., 2003, 75, 2117±2122. 153. Spatz J P, Chan V Z H, Moessmer S, Kamm F M, Plettl A, Ziemann P, Moeller M, `A combined top-down/bottom-up approach to the microscopic localization of metallic nanodots' Adv. Mater., 2002, 14, 1827±1831. 154. Glass R, Moeller M, Spatz J P, `Block copolymer micelle nanolithography' Nanotechnology, 2003, 14, 1153±1160. 155. Spatz J P, Moessmer S, Hartmann C, Moeller M, Herzog T, Krieger M, Boyen H G, Ziemann P, Kabius B, `Ordered deposition of inorganic clusters from micellar block copolymer films' Langmuir, 2000, 16, 407±415. 156. Glass R, Arnold M, Bluemmel J, Kueller A, Moeller M, Spatz J P, `Micronanostructured interfaces fabricated by the use of inorganic block copolymer micellar monolayers as negative resist for electron-beam lithography' Adv. Funct. Mater., 2003, 13, 569±575. 157. Hanarp P, Sutherland D S, Gold J, Kasemo B, `Control of nanoparticle film structure for colloidal lithography' Colloid Surf. A, 2003, 214, 23±26. 158. Frey W, Woods C K, Chilkoti A, `Ultraflat Nanosphere lithography: a new method to fabricate flat nanostructures' Adv. Mater., 2000, 12, 1515±1519. 159. Mosbach K, Ramstrom O, `The emerging technique of molecular imprinting and its future impact on biotechnology' Bio/technology, 1996, 14, 163±170. 160. Steinke J, Sherrington D, Dunkin I, `Imprinting of synthetic polymers using molecular templates' Adv. Polym. Sci., 1995, 123, 80±125. 161. Glad M, Norrlow O, Sellergren B, Sieghban N, Mosbach K, `Use of silane

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monomers for molecular imprinting and enzyme entrapment in polysiloxane-coated porous silica' J. Chromatogr., 1985, 347, 11±23. 162. Venton D L, Gudipati E, `Entrapment of enzyme using organo-functionalized polysiloxane copolymers' Biochim. Biophys. Acta, 1995, 1250, 117±125. 163. Shi H, Tsai W B, Garrison M D, Ferrati S, Ratner B D, `Template-imprinted nanostructured surfaces for protein recognition' Nature, 1999, 398, 593±597. 164. Xia Y N, Rogers J A., Paul K E, Whitesides G M, `Unconventional Methods for Fabricating and Patterning Nanostructures' Chem. Rev., 1999, 99, 1823±1848.

Part II

Measurement, monitoring and characterisation

7

Surface spectroscopies

M M E H L M A N N and G G A U G L I T Z , University of Tuebingen, Gemany

7.1

Introduction

In recent years, some focus of research has been on biochemical and chemical sensors. The combination of physical sensors (transducers) with more or less analyte-selective layers of biochemical or chemical substrates has introduced selectivity to these systems. For this reason such set-ups have to be considered as complete sensor systems containing transduction principles, the sensitive layer, the signal processing, and evaluation strategies. Out of the huge variety of transduction principles, this chapter concentrates on optical techniques which provide many possibilities of application of optical principles. As this chapter is a review based on a lecture, a large number of optical principles will be classified and a survey on sensitive layers that differ as to sensitivity, selectivity, stability, and reversibility will be discussed. To cover the wide field in part, recent review articles have been included. This chapter is divided into three main sections. In the first section different surfaces for biomolecular sensing are discussed and examples of how to characterise these surfaces are given. In the second, different optical sensing methods are shown and compared with each other. In the last section some selected applications of biomolecular interaction analysis are given. The numerous publications in the area of optical sensing allow only the citation of some related publications and reviews.

7.2

Surfaces

A very large number of bioanalytical methods deal with monitoring interactions occurring at surfaces or at least uses surface bound ligands or receptors to monitor changes in analyte concentration even within a homogeneous phase. For all these detection methods there is a need for very special and smart surface properties. Selectivity, sensitivity, stability, and reversibility are examples of such requirements for these sensor systems which must be provided in part by the surface properties. In the case of biosensors, the user expects a rather high signal-

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to-noise ratio, short response times, low limits of detection, high sensitivity, and the possibility to use sensors also in real samples, not just in lab applications. Based on ion-selective electrodes, at the beginning, doped semi-conductor material was used as a sensing system. Later on layers derived from chromatography (Baldine and Bracci, 2000) were used as rather stable layers with high reversibility and very short response times. However these layers showed a quite low selectivity and especially for the detection of gases the limit of detection stayed only in the range of ppm. (Rathgeb and Gauglitz, 2000). Another approach to introduce selectivity to these layers is to use microporous material. According to the free volume of the micropores a discrimination by size is achieved (Lehner, 1996). Combining this principle with swelling properties of selected polymer films to detect gases or liquids allows even a discrimination by molecular dimensions and partition coefficients (Dieterle et al., 2002). For many years biomolecules have been used to provide very high selectivity and sensitivity. For example, antigen/antibody interactions, DNA/DNA hybridisation, inhibition of enzymes, protein/protein interaction and many other applications in the area of membrane bound receptors and signal cascades can be found. These biomolecules should exhibit a high receptor/ligand specificity, whereas polymers or organic sensitive layers exhibit non-specific interaction at higher reversibility. For all detection methods based on heterogeneous phase interactions nonspecific binding effects are a very important point. Therefore very sophisticated surface chemistry and modifications are a basic requirement to set up an efficient biosensor. One important approach is the silanisation of glass or quartz transducers with the subsequent covalent binding of various biopolymers supplying reduced non-specific binding properties and allowing functionalisation with ligands or receptors. This silanisation step can be characterised by NMR-spectroscopy and ellipsometry (Raitza et al., 2000). Dextran hydrogels (LoÈfas and Johnsson, 1990) supply a large number of functional sites within the volume (Piehler et al., 1996). However, in many cases, especially observing protein interactions non-specifity is not reduced enough. To solve this problems polyethylenglycole with different chain lengths bound to the silanised surface can be used as a shielding polymer against nonspecific binding (Feldmann et al., 1999). Ligands can easily be immobilised by means of either amino or carboxy functions. These layers are quite resistant against non-specific binding (Piehler et al., 2000). In contrast to the dextran hydrogel, the PEG-surfaces form a kind of two-dimensional polymer brush and interaction occurs only at the surface and not in the volume. Therefore these layers have a reduced number of interaction sites. Besides these principal approaches, a large number of other ideas have been realised, e.g., the use of avidin to immobilise biotinylated biochemical molecules to the transducer surface (Birkert et al., 2000), the use of His-tags

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(Gershon and Khilko, 1995) or the immobilisation of membrane structures of lipid double layers (Tien, 1985) to the transducer to model cell walls. All these various approaches have the intention to reduce non-specific binding, allow a large number of specific binding sites, increase the stability of the layer, which is essential for regeneration strategies, and increase selectivity and sensitivity. The disadvantage of such approaches is that due to the high binding constant the system is not reversible at all. Therefore regeneration strategies have to be introduced to regain the reusability of these sensing surfaces. One major drawback of these biomolecules is that although they increase the stability of the biolayer, their stability is not comparable with that of, e.g., polysiloxane films or microporous systems (Park et al., 1999). In order to try to combine the advantages of stability and reversibility with sensitivity and selectivity a wide field of research for many years has been supramolecular structures (Lehn and Ball, 2000) and biomimetic layers (Garnier, 2000). To increase selectivity of simple chemosensors supramolecular structures such as calixarene (Dickert and Schuster, 1995) were tried first. Another approach was the use of cyclodextrins (Schurig and Grosenick, 1994) or cycloheptapeptide (Jung et al., 1996) structures. Hybridisation studies were one of the first approaches to introduce selectivity by immobilising polynucleotide or peptide sequences. In the meanwhile, peptide nucleic acids (PNA, with peptides as a backbone) (Wang, 1999) have proven to be a better complementary binding system than DNA because of less repulsion by charges and better backbone stability, thus being stable against DNases and nucleases. Another approach is the synthesis of layers of molecular imprinted polymers (Haupt and Mosbach, 2000). During a co-polymerisation process a template is used to form a cavity with certain spacings and some selective binding sites. After removing the template the generated recognition and binding site is highly selective for the template (or molecules similar to the template) as a considered analyte. Although these layers are very stable and selective they often show very slow reaction times upon analyte exposure. To increase reaction times in the socalled spreader bar technique (Mirsky et al., 1999) the imprinting process is reduced only to the surface instead of the volume. Whether these layers will show potential in the future has yet to be proven. Figure 7.1 shows various different possibilities of assay formats. The best possibility would be a homogeneous assay which allows a direct interaction between the analyte and the reagent without any influence on surface properties. Especially in the case of optical detection, the problem is that this interaction has to change somehow the optical properties of the system. These might be simple colour changes. Therefore, in the case of homogeneous assays, mostly labelled systems are used. Although radioactive labelling allows a quite low limit of detection a fluorescence label is preferable. Therefore fluorescence labelling is normally used, taking advantage of either quenching effects or (in most cases) fluorescence resonance energy transfer.

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7.1 Assays using interaction between biomolecules in homogeneous phase and/or at heterogeneous interfaces. In both cases thermodynamics (equilibrium constant) and kinetics (association and dissociation rate constants) determine the interaction. Direct assays immobilise the receptor at the surface to measure the analyte, here a binding inhibition assay is demonstrated where derivatives of the analyte or ligand to detect are immobilised. In the preincubation phase receptor and ligand are mixed in the homogeneous phase, concentration of non-blocked receptor molecules is detected via the heterogeneous phase. High numbers of interaction sites at the transducer make this process diffusion controlled, at low `loading' the kinetics at the heterogeneous phase can be measured.

In the case of measurements at the heterogeneous phase two different detection principles can be distinguished, direct optical detection without any label and detection using fluorescence labels. In the case of direct optical detection preferably large analyte molecules have to be examined. The sensitivity of any kind of optical detection can be improved by increasing the analyte mass or volume. Therefore, either a competitive test scheme (labelled competing with non-labelled analyte) or a so-called binding inhibition test scheme is used. This represents the following assay type: an analyte derivative is immobilised in the biopolymer; in a preincubation phase the analyte and the reagent are mixed together; the analyte as a ligand blocks receptor sites, which cannot react any more in this blocked state with the analyte derivative immobilised to the surface. This approach can be realised with a labelled reagent or direct optical applications.

7.3

Optical detection methods

Monitoring of effects at surfaces or interfaces using regular reflectance of light can be divided into two basic detection principles based on reflectometry or refractometry. Both use the influence of the thickness and/or refraction index of the observed layer on the phase and/or amplitude of electromagnetic radiation

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7.2 Schematic drawing of regular reflection, resulting refraction and total internal reflection for angles larger than the critical angle in the waveguide. The guided wave exhibits an evanescent field close to the waveguide, which is influenced by absorbing molecules or changes in the refractive index within the penetration depth of this evanescent field. Accordingly, the coupling of this external field to the guided field vectors within the waveguide influences the effective refractive, the transversal electric (TE-) and magnetic (TM-) differently (Hecht and Zajec, 1980).

penetrating this layer or being reflected. Different detection principles are schematically demonstrated in Fig. 7.2. In the case of refractometry, radiation guided in a wave guide is influenced by minimum changes in the refractive index or transmittance of an adjacent medium, since its evanescent field probes this medium resulting in an effective refractive index. Thus the TE (transversal electrical) and TM (transversal magnetic) modes of the wave propagating in the wave guide are influenced differently. A review of a variety of different optical detection principles based on evanescent field techniques can be found in Gauglitz (1996). Figure 7.2 demonstrates that incident radiation is reflected in part at an interface between media with different refractive indices according to the Fresnel equations. Radiation passing through the medium with high refractive index is refracted to the optical axis (Fig. 7.2, left side). Radiation passing through the interface in the opposite direction will be refracted away from the optical axis. Thus, a critical angle exists beyond which radiation will not pass through the interface out of the medium of high refractive index; it will be `totally reflected'. Accordingly, under such conditions, radiation will be guided in this medium being a fibre or, more generally, a waveguide (planar or stripe). The electric field of this guided wave causes an evanescent field existing close to the interface in the medium of lower refractive index. The evanescent field penetrates the adjacent medium just by approx. half of the wavelength of the

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guided radiation and also shows an exponential decay. Therefore, any changes in these properties influence the evanescent field and accordingly the propagation of radiation in the waveguides. The `effective refractive index' of the waveguide varies. Interferometer type waveguides like Mach-Zehnder interferometer (Ingenhoff et al., 1993) or Young interferometer (Brandenburg et al., 1992) use two different waves guided in two different waveguides. One of these waveguides is exposed to the analyte while the other is used as a reference channel. This leads to a detectable difference in phase of the two waves. For the Young interferometer the waveguide arms are not reunifed but rather image the interference pattern produced by the two open ends of the waveguide arms on a CCD (Brandenburg et al., 1992). Using both TE and TM modes allows internal referencing. This has been improved by Lukosz and Stamm (1991) in its mode beat interferometer, measuring amplitudes and phases of both polarisation states. Another device that is frequently used for monitoring changes in refractive index at surfaces is the grating coupler (Clerc and Lukosz, 1994). The grating coupler combines a waveguide layer with layers in which a grating is embedded. The grating constant is influenced by the refractive index within the adjacent medium. Similar to an interference filter, this grating condition varies with the angle of incidence or determines the preferred wavelengths. Dependent on the effective refractive index (that is strongly influenced by the adjacent layer where the interaction takes place), wavelength and angle of incidence radiation incident to the grating will be reflected or coupled in the waveguide (Nellen and Lukosz, 1993). The reflected radiation is monitored using either an angle resolved set-up or a CCD camera (avoiding mechanical parts). A special set-up uses so-called bi-diffractive couplers where two different grating constants are superimposed resulting in a different angle for the outcoupled wave compared to the direct reflected one. Another type of interrogation of the polarisation status is applied in prism couplers (Cush et al., 1993). The radiation couples out of a prism via the frustrated total internal reflection of a low refractive index layer (with a thickness of ~1000 nm) into the high refractive index waveguide. A 45ë polarisation is chosen, and TM and TE modes travel in the resonant layer (waveguide: thickness of ~100 nm), differently influenced via evanescent field by changes in the adjacent medium. Thus, the polarisation state changes in this `resonant mirror'. The angle at which this coupling occurs, the resonant angle, is, essentially, dependent upon the refractive index at the surface of the sensor. Changes in refractive index of the adjacent layer will change the resonant angle. The best researched evanescent field technique is surface plasmon resonance being reviewed in theory and application to chemo- and biosensing in many articles (Homola et al., 1999, 2002). A prism is coated at its basis by an approx. 50 nm metal film. When the energy of the incident photon electrical field is just right it can interact with the free electron constellations in the metal film. The

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incident light photons are absorbed and converted into surface plasmons resulting in a `dip' in the reflectance diagram. The resonance condition of these plasmons depends, via the evanescent wave, also on the refractive index at the surface opposite to the waveguide. Simplified, the binding of biomolecules results in the change of the refractive index on the film, which is measured as a change in reflected light. This direct optical detection method has been commercialised the longest. In recent years, additional reviews have been published regarding SPR techniques (Rich and Myszka, 2000; Van Der Merwe and Anton, 2001; Davis and Wilson, 2001; Sadana, 2001), resonant mirror (Kinning and Edwards, 2002), grating coupler (Voros et al., 2002; Kuhlmeier et al., 2003), Bragg gratings (Santos and Ferreira, 2002).

7.3.1 Reflectometry At the interface of a thin layer the incident light is partially reflected (depending on the difference of refractive indices of the two layers) whereas the other part penetrates the layer and is again partially reflected at the second interface. These two partially reflected beams can superimpose and form a characteristic interference pattern depending on the wavelength, the angle of incidence and the refractive index and the physical thickness (the so-called `optical thickness') of this layer. Changes on or at this layer (for example, due to analyte binding to this layer) result in a characteristic shift of this interference pattern that is easily detected. The principle of this simplified version of ellipsometry as shown in Fig.7.3 is called reflectometric interference spectroscopy (RIfS). It is a simple and robust technique in chemo- and biosensing as demonstrated later in the application section. In ellipsometry the polarised light is used to probe the dielectric properties of a sample. The most common application of ellipsometry is the analysis of very thin films. Through the analysis of the state of polarisation of the light that is reflected from the sample, ellipsometry can yield information about layers that are thinner than the wavelength of the light itself, down to a single atomic layer or less. Depending on what is already known about the sample, the technique can probe a range of properties including the layer thickness, refractive index, morphology, or chemical composition. The first introduction of ellipsometry was as far back as the 1940s (Azzam and Bahara, 1988; Arwin and Aspnes, 1986), and today it has many standard applications. It is mainly used in semiconductor research and fabrication to determine properties of layer stacks of thin films and the interfaces between the layers. However, ellipsometry is also becoming more interesting to researchers in other disciplines such as biosensors and medicine. These new disciplines require new techniques such as microscopic imaging. The polarisation state of the light beam can be determined in many different ways. In the nulling technique the polarising elements were rotated until the

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Surfaces and interfaces for biomaterials

7.3 The principle of reflectometric interference spectroscopy (RIfS) is based on white light interference according to the given formula. Changes in the optical thickness of the layer between the two interfaces causes a change from second reflected beam I2 to I0 2 superimposed to I1 and from a destructive interference (at a specific wavelength: broken line) to a positive constructive interference. Considering the whole interference pattern a shift is observed correlating to the amount of change in physical thickness. IR can be used as a reference. The change in physical thickness is caused either by swelling of a polymer layer (uptake of analyte) or by an affinity reaction adding biomolecules to receptor molecules at the interface.

signal at the detector is extinguished (nulled). This technique was also used in the first ellipsometers. The measuring of a signal near `zero' is one disadvantage of this nulling technique, as modern light detectors have a significantly higher noise at low signal intensities. Therefore other techniques to determine the polarisation state, e.g., phase modulated ellipsometry, have been developed. Both principles (reflectometry and refractometry) measure changes in the optical thickness of a sample. However, it has to be considered that the refractive index is a highly temperature dependent value. Thus all devices based on evanescent field techniques need a good thermostating (below 0.1 K) or need to be referenced well using a dual-channel instrument. However, in the case of reflectometry thermostating normally is not required. By chance an increase of the temperature leads to an increase of the thickness of the layer due to thermal volume expansion. This increase results in a decrease of the refractive index of the layer (the increase of the layer leads to a lower density of polarisable molecules and thus to a lower refractive index). These two effects nearly compensate each other.

7.4

Biomolecular interaction analysis

In principle, monitoring of interactions with label-free biosensors can be divided into two assay formats. First, the interaction between the immobilised ligand and

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7.4 Scheme of a simple binding experiment. Under appropriate experimental conditions the association and the dissociation rate constants, ka and kd, and also the affinity constant K can be determined.

the dissolved binding partner can be monitored directly, enabling the study of kinetics and the thermodynamics of the (heterogeneous) reaction. This direct detection is used frequently for macromolecules like proteins. An indirect assay format is often applied for the detection of interactions between a macromolecule and a low molecular weight species. In this type of assay both examined binding partners are in solution and the equilibrium concentration is quantified with the biosensor, allowing the calculation of the thermodynamic properties of the `homogeneous' reaction (Fig. 7.4).

7.4.1 Direct detection The easiest case is a simple 1:1 stoichiometric reaction according to: A‡Bˆ

ka kd

7:1

where A is the immobilised binding partner onto the transducer surface, B the soluble binding partner, and ka and kd are the association and the dissociation rate constants, respectively. Assuming a constant concentration of the soluble

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binding partner (for example, in a constant flow device) the binding reaction can be described as a pseudo-first-order reaction according to the Langmuirian theory: dÿ=dt ˆ ka C0 …ÿmax ÿ ÿ† ÿ kd ÿ

7:2

where ÿ is the surface coverage of captured ligate, C0 is the solution concentration of ligate, and ÿmax is the total binding capacity of the surface. Assumptions made for the validity of this model are: the binding event is controlled only by the intrinsic kinetic rate constants ka and kd and the diffusion of the soluble binding partner to or from the surface is not rate-limiting, the concentration of the ligate C0 is constant; the interaction of the biomolecules occurs in 1:1 stoichiometric ratio. Either by applying a linear transformation to eqn 7.2 according to Karlsson et al. (1991) or by an exponential curve fitting as proposed by O'Shannessy et al. (1993) the corresponding values for ka and kd can be obtained. The dissociation rate constant kd can also be calculated from the dissociation of the bound ligate from the surface during the dissociation (wash out) phase by replacing the ligate solution by buffer. The dissociation can be described by the first-order equation: dÿ=dt ˆ ÿkd ÿ

7:3

Even if the methods of linear transformation are simple and practical to use they imply the disadvantage of error propagation by the transformation and suppress the fact that the data are often non-linear. Therefore it is clear that this kind of analysis gave rise to criticism (O'Shannessy, 1994). The benefit of an exponential data analysis is that the errors of the kinetic rate constants directly reflect errors in the primary data and the curve fitting is permanent under the control of the operator. Moreover the values for ÿeq (the extrapolated equilibrium response) and ÿmax are fit which is possible even if the binding curve does not reach the equilibrium within the measurement time. The third method to derive kinetic rate constants is also an iterative curve fitting routine called global analysis (Roden and Myszka, 1996). This method is applied simultaneously to complete sets of binding curves at varied conditions. Using a reference channel the data can be corrected for drift, non-specific binding and refractive index effects. A comparison of the three analysis routines, linearization, integrated rate equation and numerical integration is given in Morton et al. (1995). A major problem of all these analysis methods is the possible difference between the assumption of a first-order reaction limited only by reaction kinetics to the real experimental data. Significant deviations from ideal conditions will lead to wrong results. For example Nieba et al. (1996) found large differences between dissociation constants derived from kinetic rate constants measured at the surface and determined in the solution.

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Possible differences between experiment and model may be a non-negligible influence of mass transport to the surface, an underestimation of the dissociation constant due to rebinding effects or a more complex binding mechanism that could not be described sufficient by eqn 7.1. A detailed description of the influence of mass transport on the data can be found in Glaser (1993). The most important parameters are the product ka0 g and the Onsager coefficent of mass transport Lm, with ka0 as the observed associate rate constant and g the surface concentration of free binding sites. The Onsager coefficient is according to SjoÈlander and Urbaniczky (1991) dependent on the geometrical dimensions of the flowcell r 2 3 D f Lm ˆ km ˆ 0:98
7:4 h2 bl with the diffusion coefficient D, the flowrate f, the channel height h and its width b and the distance l form flow cell entrance. According to Glaser three cases can be distinguished: 1. 2. 3.

ka0 g  Lm , with no mass transport influence on the kinetics. ka0 g and Lm are in the same order of magnitude. Then the true rate constants ka and kd are underestimated up to several orders of magnitude. ka0 g  Lm with full mass transport control. No kinetic information can be gained.

Methods for data evaluation of mass transport limited binding kinetics are suggested in several publications (Schuck, 1996; Myszka et al., 1996; Schuck and Minton, 1996). For data analysis, very often the association phase is split into the mass transport limited phase and the kinetically limited phase. This can be carried out by the determination of the linear part of a dCt/dt versus Ct plot. However, especially for high association constants this method underestimates the values of ka. Schuck and Minton (1996) suggested a two-compartment model taking mass transport into account and considering the transport only from the bulk compartment to the surface. However, the best way to resolve this problem is to reduce the mass transport effects by special experimental conditions. According to eqn 7.4 the Onsager coefficient can be influenced by using appropriate flowrates and cell geometries, and the reduction of the surface concentration of free binding sites g will reduce the influence of mass transport (the apparent rate constant ka0 is a fixed constant for a given biochemical system and cannot be changed). On the other hand the reduction of the binding capacity results in lower signals. Thus a high sensitivity and low S/N ratios of the biosensor equipment are required. These parameters determine the upper limit of the range of association rate constants amenable to study by optical biosensors (Hall et al., 1996). Another reason for differences between experiment and the simple model according to eqn 7.1 is the so-called rebinding. This rebinding results in an

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Surfaces and interfaces for biomaterials

underestimation of the dissociation rate constant kd. The probability for the rebinding to the surface increases with the concentration of free binding sites (Shoup and Szabo, 1982). This influence can also be reduced by lowering the surface concentration of free binding sites g, or by adding a competitive binder in high concentration to the washing buffer, to prevent dissociated molecules rebinding to the surface. The (heterogeneous) thermodynamic constant KA can be calculated either from the kinetic rate constants ka and kd or by evaluating the concentration dependent equilibrium binding signals. For 1:1 reactions the Langmuir adsorption isotherm is: ÿeq ˆ ÿmax

C0 1 ‡ C0 KA

7:5

A Scatchard Plot ÿeq =C0 versus ÿeq results in a straight line with the slope ÿKA.

7.4.2 Indirect detection The constants measured at the surface often differ from those in the solution. Therefore a different approach to determine the homogeneous association constant KA has been developed using the sensor surface to probe for the concentration of unbound analyte molecules in solution. The principle is given in Fig. 7.5. An equilibrated mixture of analyte and the corresponding binding partner (ligate) is injected over the sensor surface where an analyte-derivative is immobilised on it. From the initial binding rate the concentration of unbound ligate in solution can be determined (Karlsson et al., 1994; Edwards and Leatherbarrow, 1997). According to eqn 7.2, for the beginning of the binding (ÿ ˆ 0 pg/mm2): dÿ=dt ˆ ka C ÿmax

7:6

the signal is directly proportional to the concentration of unbound ligate C. If the reaction is under a pure mass transport control the initial binding rate can be described according to Fick's first law: dÿ 7:7 ˆ km
C dt with km as the diffusion coefficient. Therefore in both cases the initial binding rate is directly proportional to the concentration of unbound ligate in solution. A binding rate proportional to concentration of free ligate C is always given if the amount of bound ligate does not exceed 10% of the maximum binding capacity, therefore high binding capacities of the surface are desirable (Piehler et al., 1997a).

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7.5 Probing of the equilibrium concentration of unbound ligand with the surface-bound ligand.

For a known initial concentration of ligate C0 the dissociation constant can be derived from: c
a …C0 ÿ x†
…A0 ÿ x† ˆ 7:8 x x with c and a as the equilibrium concentrations of ligate and analyte, and C0 and A0 as the corresponding initial concentrations, and x as the equilibrium concentration of formed complex. Equation 7.8 can be rearranged to: s C0 ÿ A0 KD …C0 ‡ A0 ‡ KD †2 ‡ ÿ C 0 A0 cˆ 7:9 2 4 KD ˆ

The resulting secondary plot (titration curve) for different concentrations of analyte C0 is shown in Fig. 7.5. By fitting eqn 7.9 to the titration curve the dissociation constant KD can be determined. This kind of assay is often denoted as a competitive format, but this term implies that a competition between ligate,

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Surfaces and interfaces for biomaterials

(immobilised) analyte derivative and analyte (dissolved) occurs. Due to the short time of interaction in a flow system between the surface and the solution of ligate and analyte no equilibrium for these three species can be reached (in fact the equilibrium in the solution is not affected by the surface if dissociation of the ligate-analyte complex is not too high) and therefore the binding rate depends only on the concentration of free ligate. For this reason the term inhibition test is a better choice. The advantages of this assay type are that due to the indirect assay format also small analytes can be investigated, and that the effects caused by the immobilisation (rebinding, mass transport) do not affect the results. Also many different analytes can be examined using the same surface. Recently, even a new detection principle has been developed. A time resolved observation of the interaction process during the incubation step allows even the determination of the on and off-rates (ka and kd) for the homogeneous reaction (between ligand and dissolved analyte). In this case, the surface is used only to determine the concentration of unbound ligand in solution. Due to steric hindrance, for example, these on and off-rates can be quite different compared to the on and off-rates for the heterogeneous reaction (between ligand and surface bound analyte).

7.5

Conclusion

Optical sensors have proven in the past to be either very simple and cost-effective devices or to allow rather sophisticated multisensor applications. As a large number of different optical principles exist, in principle many of these methods can be applied to a huge number of applications. It has to be noted that in most cases the choice of an appropriate assay design, and therefore the surfaces and transducers used are linked together very closely and cannot be considered separately. It becomes evident that none of the afore-mentioned various sensing principles, and also the unmentioned electrochemical or mass sensitive devices, is superior in general. Their feasibility is strongly dependent upon the application they are used for. This became obvious when comparing various refractometric and reflectometric methods in the same biomolecular interaction study using antibodies and antigens, and setting up the surface chemistry by the same person. In the case of both studies, the limits of detection ranged for all methods examined within one order of magnitude. The discrimination was achieved only at the cost of expenditure on apparatus and by the sophistication of the fluidics used (HaÈnel and Gauglitz, 2002; Piehler et al., 1997b). The essential result of these considerations is certainly that research in the area of sensing requires interdisciplinary understanding of the detection principles, of the sensitive layer, of the kinetics and thermodynamics of interaction processes and of the fluidics. Thus fundamental research has to characterise these layers and the interaction processes to improve the understanding which is the prerequisite of any optimisation approach.

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197

References

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Sens Actuator, B54, 3. Homola J, Yee S, Myszka D (2002), in Optical Biosensors: Present and Future, eds Ligler F S, Rowe T, Chris A, Elsevier Science BV, Amsterdam, 207. Ingenhoff J, Drapp B, Gauglitz G (1993), `Biosensors using Integrated Optical Devices', Fresenius J. Anal. Chem., 346, 580±583. Jung G, Hofstetter H, Feiertag S, Stoll D, Hofstetter O, WiesmuÈller K-H, Schurig V (1996), `Cyclopeptide libraries as new chiral selectors in capillary electrophoresis', Angew Chem Int Ed Engl, 35, 2148. Karlsson R, Michaelsson A, Mattsson L (1991), `Kinetic analysis of monoclonal antibody-antigen interactions with a new biosensor based analytical system', J Immunol Methods 145, 229-240. Karlsson R, Roos H, FaÈgerstam L, Persson B (1994), `Kinetic and concentration analysis using BIA technology', Methods: Companion Methods Enzymo, 6, 99±110. Kinning T, Edwards P (2002), in Optical Biosensors: Present and Future, eds Ligler F S, Rowe T, Chris A, Elsevier Science BV Amsterdam, 253. Kuhlmeier D, Rodda E, Kolarik L O, Furlong D N, Bilitewski U (2003), `Application of atomic force microscopy and grating coupler for the characterization of biosensor surfaces', Biosens Bioelectron, 18, 925. Lehn J M, Ball P (2000), `Supramolecular chemistry', in New Chemistry, ed. N Hall, Cambridge University Press, 300. Lehner M D (1996), Macromolecular Chemistry: A Textbook for Chemists, Physicists, Material Scientists, and Process Technicians, BirkhaÈuser Verlag, Basel. LoÈfas L, Johnsson B (1990), `A novel hydrogel matrix on gold surfaces in surface plasmon resonance sensors for fast and efficient covalent immobilization of ligands', J Chem Soc Chem Commun, 1526. Lukosz W, Stamm C (1991), `Integrated optical interferometer as relative humidity sensor and differential refractometer', Sens Actuator, A25, 185. Mirsky V M, Hirsch T, Piletsky S, Wolfbeis O S (1999), `A spreader-bar approach to molecular architecture: formation of stable artificial chemoreceptors', Angew Chem Int Ed Engl, 38, 1108. Morton T A, Myszka D G, Chaiken I M (1995), `Interpreting complex binding kinetics from optical biosensors: a comparison of analysis by linearization, the integrated rate equation, and numerical integration', Anal Biochem, 227, 176±185. Myszka D G, Morton T A, Doyle M L, Chaiken I M (1996), `Kinetic analysis of a protein antigen-antibody interaction limited by mass transport on an optical biosensor', Biophys Chem, 64, 127±137. Nellen P M, Lukosz W (1993), `Model experiments with integrated optical input grating couplers as direct immunosensors', Biosens. Bioelectr., 6(6), 517±25. Nieba L, Krebber A, PluÈckthun A (1996), `Competition BIAcore for measuring true affinities: large differences from values determined from binding kinetics', Anal Biochem, 234, 155±165. O'Shannessy D J (1994), `Determination of kinetic rate and equilibrium binding constants for macromolecular interactions: a critique of the surface plasmon resonance literature', Curr Opin Biotechnol, 5, 65±71. O'Shannessy D J, Brigham-Burke M, Soneson K K, Hensley P, Brooks I (1993), `Determination of rate and equilibrium binding constants for macromolecular interactions using surface plasmon resonance: use of nonlinear least squares analysis methods', Anal Biochem, 212, 457±468.

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Park J, Groves W A, Zellers E T (1999), `Vapor recognition with small arrays of polymer-coated microsensors. A comprehensive analysis', Anal Chem, 71, 3877. Piehler J, Brecht A, Geckeler K E, Gauglitz G (1996), `Surface modification for direct immunoprobes', Biosens Bioelectron, 11, 579. Piehler J, Brecht A, Giersch T, Hock B, Gauglitz G (1997a), `Assessment of affinity constants by rapid solid phase detection of equilibrium binding in a flow system', J. of Immunol. Meth., 201(2), 189±206. Piehler J, Brandenburg A, Brecht A, Wagner E, Gauglitz G (1997b), `Characterization of grating couplers for affinity-based pesticide sensing', Appl. Opt., 36, 6554. Piehler J, Brecht A, Valiokas R, Lidberg B, Gauglitz G (2000), `A high-density poly(ethylene glycol) polymer brush for immobilization on glass-type surfaces', Biosens Bioelectron, 15, 473. Raitza M, Herold M, Ellwanger A, Gauglitz G, Albert K (2000), `Solid-state NMR and ellipsometric investigations of C30 chains bonded to SiO2 surfaces', Macromol Chem Phys, 201, 825. Rathgeb F, Gauglitz G (2000), `Optical gas sensors in analytical chemistry: applications and trends and general comments', in Encyclopedia of Analytical Chemistry, ed. R A Meyers, John Wiley & Sons Ltd, Chichester, 2189. Rich R L, Myszka D G (2000), `Survey of the 1999 surface plasmon resonance biosensor literature', J Mol Recogn, 13, 388. Roden L D, Myszka D G (1996), `Global analysis of a macromolecular interaction measured on BIAcore', Biochem Biophys Res Com, 225, 1073-1077. Sadana A (2001), `Kinetic analysis for analyte-receptor binding and dissociation in biosensor applications: a fractal analysis', Biotech Genetic Eng Rev, 18, 29. Santos J L, Ferreira L A (2002), `Fibre Bragg grating interrogation techniques', in Handbook of Optical Fibre Sensing Technology, John Wiley & Sons, Chichester, 379. Schuck P (1996), `Kinetics of ligand binding to receptor immobilized in a polymer matrix, as detected with an evanescent wave biosensor. I. A computer simulation of the influence of mass transport', Biophysical J, 70, 1230±1249. Schuck P, Minton A P (1996), `Analysis of mass transport-limited binding kinetics in evanescent wave biosensors', Anal Biochem 240: 262±272. Schurig V, Grosenick H (1994), `Preparative enantiomer separation of enflurane and isoflurane by inclusion gas chromatography', J Chromatography, A666 617. Shoup D, Szabo A (1982), `Role of diffusion in ligand binding to macromolecules and cell-bound receptors', Biophys J, 40, 33±39. SjoÈlander S, Urbaniczky C (1991), `Integrated fluid handling system for biomolecular interaction analysis', Anal Chem, 63, 2338±2345. Tien H T (1985), `Planar bilayer lipid membranes', Prog Surf Sc, 19, 169. Van Der Merwe, Anton P (2001), `Surface plasmon resonance', in Protein-Ligand Interactions: Hydrodynamics and Calorimetry, Oxford University Press, 137. Voros J, Ramsden J J, Scucs G, Szendro I, De Paul S M, Textor M, Spencer N D (2002), `Optical grating coupler biosensors', Biomaterials, 23, 17, 3699. Wang J (1999), `PNA biosensors for nucleic acid detection', Curr Issue Molec Biol, 1(2), 117.

8

Surface microscopies C Z I E G L E R , University of Kaiserslautern, and Institut fuÈr OberflaÈchen- und Schichtanalytik GmbH, Kaiserslautern, Germany

8.1

Introduction

Biomaterials must meet the demands of materials science on various length scales as well as clinical requirements. In particular, the interface is important for biomaterials, as it defines the interaction with the environment,1,2 as described elsewhere in this book. Interface structures are in dimensions of a few nanometers for proteins up to the micrometer range for cells in the field of biomaterials. In Fig. 8.1 an overview of the interfaces on different scales is presented to show the parameters that can be of interest. Macroscopically, the shape of an implant, the structure, and its mechanical stability are important. The microscopic level is determined respectively by morphology (i.e. domain structure, the presence of ionic groups and chemical composition respectively, including further modification), topography (i.e. surface roughness, planarity, feature dimensions), hardness and elasticity (Young's modulus). These characteristics determine other features like wetting behavior and interaction forces (inter- and intramolecular) including cell-surface or cell-cell interactions. Various degrees of information about these properties can be obtained using different analysis methods including microscopic and spectroscopic methods. Mechanical, biochemical, and optical properties depend mainly on the topography and chemistry of the surface. Topography may be defined as size, shape, distribution, and the hierarchy of surface features. These features are either discrete (holes and peaks) or continuous (furrows and ridges) in random (statistical), fractal (self-similar on different length scales) or periodic distribution across the surface. Rough surfaces exhibit a larger surface area, and more contact points for biological molecules and cells. The reason for larger entities like cells to adhere to the surface and how they arrange their layer growth is dependent on the size and distribution of the surface features. The chemical properties of the interface between biomaterial and body environment determine the interaction of the surface with water molecules, ions, biological macromolecules and cells. Surface reactivity depends on chemical composition,

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8.1 Different relevant length scales in biomaterials research demonstrated for the example implant. The biomaterial-tissue interaction is based on molecular events, which affect meso- and macroscopic material properties. On larger scales additional (e.g. cellular, mechanical) effects resulting from combining individual units (molecules, cells, crystallites) to a more extended ensemble arise.1

the production process and pretreatment before use of the materials, and is essential for fixation, growth and proliferation of tissue and bone cells. Cell adherence via membrane receptors and adhesion proteins is influenced by active surface groups (generated, e.g., by oxidation of metals or coating of materials with specially designed polymer layers) which regulate the adsorption of the anchoring protein layer. The functionalization of the substrate additionally controls wetting behavior where particularly hydrophilic surfaces show enhanced cell adhesion. The latter is due to the fact that systems like cells

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and hydrophilic surfaces reach a thermodynamically favorable state by combining their high energetic surfaces. In addition, even hydrophobic interfaces like gold can get more hydrophilic by protein adsorption,3 which may help to increase biocompatibility. Polarizability and charging of the interface affect electrostatic interactions between charged species and the surface. Charge screening and complexation by multivalent ions present in all types of body fluids allow attraction between molecular and surface groups of like charge, and stabilize biological layers. At conductive interfaces, electrochemical reactions with charge transfer between the electrolyte solution and the substrate may occur, which interfere with cellular metabolism and the conformation of adsorbed adhesion proteins. This can set free toxic substances and cause allergic and inflammatory body reactions. Topographical as well as chemical effects play a role in the tribological properties of biomaterials. Tribology describes the behavior of interfaces in motion, and becomes important when implants are designed to support body movement. In this case, friction, wear, and lubrication of implant and body respectively, at different moving implant elements, influence the function and longevity of the applied biomaterial. Interfacial friction depends on the type and strength of interaction forces, the clasping of surface elevations and troughs, and the force transfer properties of intermediate fluids. Hardness, cohesion, and adhesion of individual surfaces in contact determine to what extent materials are worn down by abrasion (e.g. scratch and particle formation), adhesion (e.g. welding processes) and surface fatigue (e.g. crack formation). Liquid films in the crevices between two hard surfaces can reduce friction and wear. The effect of such lubricants is influenced by surface chemistry and separation, as well as by the viscoelastic, and hence force transferring, properties of the fluids themselves. This short discussion of biomaterials in relation to surface properties shows that the interplay between implant and body environment can be very complex, and includes a large variety of parameters from materials science, biology, and medicine, explained in more detail elsewhere.4 The need for surface analysis, in particular with nanometer resolution, arises on the one hand from the importance of the interface between natural and artificial materials, and on the other hand from the high dependence of macroscopic behavior on the microscopic features of the biomaterial. The following sections will show the different capabilities of today's microscopic techniques, and the criteria by which these methods could be used for certain problem-solving strategies.

8.1.1 Different concepts of imaging surfaces Imaging techniques can be categorized under different concepts. In most of the methods, probes such as beams of electrons, ions, or photons will come into

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8.2 Schematic presentation of typical arrangements to study solid surfaces.

contact with the sample surface. After interaction, either the beam of these primary probes will be analyzed for its changed properties, or secondary probes which are released from the sample surface are detected (Fig. 8.2). There are two principal types of techniques to obtain a real image of the sample. For the first one (Fig. 8.3) the probe beam has to have a small diameter and the sample surface can then be scanned point by point. The software then constructs an image from these pixels. The classical type of such an imaging instrument is the scanning electron microscope (SEM). A particular form of scanning techniques are the so-called scanning probe methods such as scanning tunneling microscopy (STM), scanning force microscopy (SFM), and scanning nearfield

8.3 Scanning techniques as one possible measuring principle to characterize geometric structures.

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optical microscopy (SNOM). Alternatively, the surface is 'illuminated` by the whole beam, and through the optical path of the instrument, a direct image can be obtained. Important techniques are the optical microscope and the transmission electron microscope (TEM). Diffraction techniques which give indirect information of the geometric structure of highly ordered materials by determining the reciprocal lattice structure will not be discussed here because biomaterials are usually not single crystalline materials.

8.1.2 Imaging parameters and requirements of the different methods To choose the right microscopic technique, one has to evaluate what information is required from the sample surface. Important questions to be answered are: · What is the required lateral resolution to see the effect? · Do I only need a topographic image of the surface, or do I need additional information, such as elemental composition, mechanical properties, or optical properties with microscopic resolution? · How surface sensitive do I need to be? Is this property present only at the surface or throughout the bulk? Do I need information on the depth to which this property occurs, or is it important only to see whether it is present or not? · Are there requirements for the sample environment, i.e., do I expect changes in properties if the sample is put into vacuum or is covered with metal etc.? Lateral resolution The lateral resolution of scanning techniques is determined by the larger diameter of the following two effects: probe beam diameter and surface area in which the analyzed information is generated. To show the difference, the interaction of electrons with a sample surface is shown in Fig. 8.4.5 The probe beam of primary electrons has a typical diameter of a few nanometers. By impacting on the surface, the primary electrons can be reflected (reflected or backscattered electrons, RE), create secondary electrons (SE) of different origin or produce X-rays. The two latter points will be explained in more detail in section 8.3. From Fig. 8.2 one can see that the SEd, which are directly produced by the primary electrons, come from an area of about the beam radius, i.e., typically 1±10 nm. The RE, the SEid, which are indirectly produced by the RE, and a special type of SE, the Auger electrons (see section 8.3) come from an area of about 1 m. Therefore the lateral resolution by imaging REs and SEid (for practical solutions see section 8.3) is in the micrometer range, whereas by imaging SEd a resolution of 10 nm or better (down to 1 nm) is possible.

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8.4 Backscattered (RE) and secondary electrons (SE) in an electron microscope. For details see text. KL are the characteristic X-rays, zSE the information depth of the secondary electrons, zA that of the Auger electrons, zRE of the backscattered electrons.5

In the far field, i.e., if the detector is far away from the sample if compared to the wavelength, the resolution of direct imaging techniques is limited mainly by the wavelength of the utilized probes because of diffraction. According to Abbe, the resolution of an imaging system depends on the wavelength , the diffraction index n of the volume between sample and objective, and the angle  under which the light passes through the objective:  8:1 n
sin  The term n
sin  is called the numerical aperture. It can be varied only to a small extent, so the resolution of light microscopy with typical wavelengths around 500 nm is limited to 200 nm. Electrons can have wavelengths in the Angstrom region, therefore atomic resolution is possible in principle. In electron microscopy, the glass lenses have to be replaced by electrical and magnetic lenses, which deflect the electron beam in an analagous way to glass lenses with A ˆ 0:61

x

x (x)

TEM

STM SFM Optical methods

SEM

Atomic resolution

x x

Height profiles x (by Auger or by EDX) x (by EELS or by EDX)

Elemental composition

Table 8.1 Measurable parameters of important imaging techniques

x (by labeling)

Specific biological groups

x

Hardness

x

Elasticity

x

Friction

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light. Because of the aberration of these electromagnetic lenses, the resolution in electron microscopy is limited to 0.2 nm. In the so-called near field, the classical diffraction limit is no longer valid (see section 8.5). Atomic resolution is therefore theoretically possible by all scanning probe techniques. However, only STM gives easily atomically resolved images; SFM shows atomically resolved structures usually only under ultrahigh-vacuum (UHV) conditions and at low temperatures, and so far no atomically resolved SNOM images have been observed due to technical problems (compare section 8.5). Additional information As outlined in section 8.2.1, there are many different properties which are of interest to fully characterize a biomaterial. Therefore, it is important to know which information can be obtained by which method. This is summarized in Table 8.1, and will be explained briefly in the respective sections below. Surface sensitivity Information depth is determined by the so-called mean free path of the primary and/or secondary probes. The mean free path is defined as the mean distance the probe can travel before it undergoes a scattering event. Scattering can result in energy and/or momentum change of the probe. Because interaction of charged particles such as electrons and ions is much larger with the atoms of the sample surface, their mean free paths are distinctly smaller than those of photons. Therefore, surface sensitivity is obtained by taking electrons or ions as primary and/or secondary probes. For the methods discussed here this is the case for SEM, TEM, and STM. However, SFM is only sensitive to the surface, because interaction forces are repulsive on a short distance scale, and therefore prevent the massive cantilever (the probe) entering at least hard surfaces (but compare section 8.4). In many cases one is interested in ascertaining depth profiles. This can be achieved by sputtering (ion etching) the surface with ions, and removing the sample layer by layer while measuring. Because the sputter process changes the sample structure, depth profiles are not combined with microscopic techniques and will therefore not be covered here. Requirements of the techniques Techniques utilizing probes with a small mean free path are surface sensitive on the one hand, but need high or even ultra high vacuum conditions in the measurement chamber, because the probes have to travel from the source to the sample, and from there to the detector without scattering with gas particles. SEM

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8.5 Finite tips cannot enter small crevices on a sample surface.

and TEM cannot therefore be used if the samples have to be studied under atmospheric or liquid conditions. (For the exception of the environmental (E-) SEM compare section 8.3.) Charged particles often produce surface charging, therefore either conductive samples are required, samples have to be made conductive (e.g. by metal evaporation), or charge compensation (if possible) has to be applied. Optical methods and SFM do not suffer from these restrictions. STM needs conductive samples, but can be applied in air and in liquid because the electrons tunnel between the tip and sample surface at a very small distance, and there is no scattering possible. All scanning probe techniques are operated with thin solid beams which come into mechanical contact with the sample surface. Because these beams have a finite diameter, they cannot enter very narrow crevices (Fig. 8.5). Very rough samples are therefore not easily imaged.

8.2

Electron microscopies

In electron microscopes, electrons are used as primary probes. Due to the small mean free paths of charged particles in air, the microscope has to be held under high vacuum conditions. In the sample, a variety of different interactions between the probe electrons and the sample atoms can take place: · transmission of primary electrons; · secondary electron emission; · Auger-electron emission, a special type of secondary electron emission which involves three electrons. Firstly, an electron hole of high binding energy (produced by the incident electron beam) is filled by an electron of lower binding energy to reduce the overall energy of the system. The energy of this process is either emitted as X-rays (characteristic X-ray emission), or used to emit a third electron of low binding energy, the Auger electron. The energy of

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·

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this electron is characteristic for the element, hence a chemical analysis can be made. electron backscattering; absorption of electrons; characteristic X-ray emission (see above), also characteristic for a certain element. Differences with Auger electrons are that X-ray emission is more efficient for higher atomic numbers and that the information with X-rays stems from 1 m depth whereas Auger electrons show an information depth of a few 10 nm. emission of light.

One big disadvantage of electrons as a probe is charging of the sample during electron bombardment. In principle, only conducting samples can be imaged. In a special case non-conducting samples can be imaged if the number of the absorbed electrons is equal to the number of secondary electrons leaving the sample. In all other cases, the sample has to be coated by metal, or a conducting replica has to be made. Both methods may give artefacts. Furthermore, the samples have to be stable under electron irradiation. There are two approaches to electron-beam use in microscopy. The first is the transmission electron microscope realized by Ernst Ruska in 1933.6,7 The TEM can be compared with a transmitted-light-microscope: electrons pass through a thin sample8 and can be detected at a fluorescence screen. Because of the strong absorption and scattering processes which take place, if free electrons interact with matter, only very thin samples in the range of 10 nm to 100 nm thickness can be passed by electrons. Sample preparation is therefore difficult and timeconsuming and can produce artefacts. As in light microscopy there are different mapping modes:9 1.

2.

3.

Bright-field-image: the image is made of electrons which directly pass through the sample. The image plane of the objective is mapped by the projector lens. The contrast is provided by the loss of electrons by scattering and absorption on their transfer through the sample. This depends on the density of the sample and the atomic number of the sample components. Diffraction image: this also uses electrons which pass directly through the sample, but in this mode the focal plane of the objective is imaged by the projector lens. By this one gets a diffraction image which yields information on the crystal structure of the sample in a similar way to an X-ray-Lauediffraction image. Dark-field-image: electrons which scatter inside a sample are imaged on the screen, and the electrons which pass through the sample in a direct way are filtered out.

Additionally, energy-filtering transmission electron microscopy11 or TEM with attached energy dispersive X-ray analysis (EDX) is used which gives the elemental composition.

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The second approach to the use of electrons as a probe for imaging is the scanning electron microscope (SEM),12 which was invented by Knoll and v. Ardenne13 in the 1930s. The first realization of an image on a surface was described by Knoll und Thiele.14 This microscope scans a very well focused electron beam over the sample surface. The electrons which leave the sample surface are accelerated by a voltage to the detector, for example an EverhartThornley-Detector,9 which transforms the electrons into an electrical signal. This is used to control intensity in a cathode-ray-oscillograph, while the same signal that scans the electron beam in SEM is used as a deflection signal in the oscillograph. So an image of the electrons leaving the sample is obtained. Resolution in SEM is estimated by the area where the electrons leave the sample. The dimensions of this area depend on the spot radius of the electron beam on the sample surface and the scattering processes inside the sample. In SEM there are several kinds of secondary electrons which can be used for image formation: 1. 2. 3. 4. 5.

Secondary electrons which are produced by primary electrons near the surface with enough energy to pass through the surface barrier. These produce the highest resolution, down to 1 nm. Secondary electrons which are produced by backscattered electrons. These produce the lowest resolution, around 100 nm. Auger electrons, which have to be detected with a special energy resolving detector (e.g., a cylindrical mirror analyzer), and give a resolution of 10 nm. Secondary electrons which are produced from backscattered electrons out of the pole shoes of the last lens. Secondary electrons produced by primary electrons which impact the aperture and produce SEs there.

The last two cases only increase noise without any further information on the sample. Backscattered electrons are too fast to reach the secondary electron detector which is placed off-axis. They can, however, also be used for imaging with a different (in-axis) detector. Backscattered electrons give high material contrast because their number depends on nuclear charge and hence atomic number. The resolution obtained is around 100 nm. The resolution of SEM depends on how well the electron beam is focused on the sample. Typical resolutions are of the order of several nm. The depth of focus can reach up to 1 mm, if compared to only 100±1000 nm in a light microscope. A review of spatially resolved spectroscopies in electron microscopy for the study of nanostructures of different metals, semiconductors, and biomaterials is given in ref. 15. To overcome the limitation of a vacuum for biological samples and/or of charging of non-conducting samples, a so-called environmental SEM (E-SEM) was invented. In this a higher partial pressure can be held in the sample chamber (up to 10-4 bar) which is differentially pumped against the rest of the

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microscope. The ionization of the background gas helps to avoid charging. Also, the higher partial pressure of, e.g., water, keeps the biological structure almost in its native state, making this method a valuable tool for imaging biomaterials.

8.3

Scanning probe microscopies

Scanning probe techniques are a valuable tool since their invention in the 1980s.16,17 In these experimental methods, distance dependent interactions like tunneling current, force, or light transmission between a sharp needle (`tip' or `probe') in close proximity to a surface (`sample') are utilized to produce an image of the sample. Two principal measurement modes were implemented: (i) to maintain a constant height of the probe above the sample while measuring the interaction change reflecting surface topography (`constant height mode'), and (ii) to maintain constant interaction while adjusting the height with the feedback signal reflecting surface topography (`constant interaction mode') (Fig. 8.6). Both modes offer the possibility of characterizing surfaces down to the atomic scale in a great variety of environments from ultra high vacuum to aqueous solutions. It is also possible to characterize time dependent reactions like crystallization and corrosion processes, as this can be done continuously and hence on-line. To scan a probe over a surface in the desired way, while reacting to the topography of the sample, positioning tools with spatial resolution in the 0.1 nm-regime are necessary. The latter is fulfilled by piezoelectric actuators

8.6 Different modes of operating scanning probe methods (a) constant height mode, (b) constant interaction mode.

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made of ceramics like PZT (lead zirconate titanate) or PMN (lead magnesium niobate), which in different directions can be extended by less than the size of one crystal unit cell. With a suitable detection unit to measure small interaction changes connected to a distance feedback circuit and digital data representation, an image of the surface can be portrayed. Experimental adaptation and extensions based on the interaction mechanisms originally used for imaging purposes lead the way to monitoring more complex features than topography, as described below. A readable overview covering many scanning probe techniques is available.18,19 In scanning tunneling microscopy (STM), the current from electrons tunneling between a conductive wire (preferably heavy metals like tungsten, platinum or iridium) with an atomically sharp tip and a (semi-)conductive surface across vacuum is measured (Fig. 8.7). The tunneling current I decays exponentially with increasing distance s between tip and surface: p ÿ p
eff 8:2 I V exp ÿk eff s s with V as the applied (Bias-)voltage, eff the effective work function, and k a constant. I additionally depends on the local densities of electronic states of the tunneling partners: X f …Ei †‰1 ÿ f …Ef †ŠR2fi …Ef ÿ Ei † 8:3 I fi

Ei,f is the energy of the initial (i) or final (f) state of the tunneling event, f(E) the Fermi function giving the probability of occupation of the energy E, and Rfi the tunneling matrix element. The delta-function stands for the energy conservation of the process. Because of its strong distance behavior, only the atom at the very end of the tip and the nearest surface atom are involved within the tunneling event. A slight change in distance, according to progression along rows of

8.7 Schematic drawing of the STM principle.

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surface atoms, alters the tunneling current in a measurable manner, so that true atomic resolution can be achieved. Experimentally, a tunneling current is obtained by securing a tip to surface distance of less than a few tenths of a nanometer under constant bias voltage. The sign of the bias determines the direction of the tunneling current, which means that unoccupied electronic surface states are probed with occupied electronic tip states or vice versa. Scanning the surface row by row either at constant height or constant current (see above) reveals surface topography. Additionally, STM can be applied to probe electronic structures. Modulating the tip-surface distance and measuring the change of the tunneling current at constant bias allows extraction of the local workfunction U of the sample. Modulating the bias and measuring corresponding current changes at a constant tip-surface distance allows extraction of the electronic states around the Fermi level of the sample (STS, scanning tunneling spectroscopy). This can lead to chemical information about the surface, but for more detailed information, electronic core levels have to be sensed which is not possible with STM (see ref. 20 and references therein). Scanning electrochemical microscopy (SECM) measures highly localized electrochemical currents associated with charge transfer reactions on metallic sample surfaces under a liquid environment.21,22 In macroscopic measurements, it can be compared with cyclic voltammetry. The reactions can occur in a fourelectrode electrochemical cell under bipotentiostatic control. There are two pathways for image production. Electron tunneling and electrochemical reactions via an electrolyte bridge occur according to the applied voltage; this can be used for the detection of localized electrochemical reactions at surfaces. It can also be used for microstructures of biomaterials like titanium.23 In addition, SECM is capable of probing the kinetics of solution reactions, adsorption phenomena and monitoring heterogeneous electron transfer kinetics associated with processes at conducting surfaces.24 The invention of scanning force microscopy (SFM) was a breakthrough for these techniques, as it became possible to image non-conducting substrates with a resolution of 0.2 nm laterally and 0.001 nm vertically. It does not require a specimen to be metal coated or stained. Non-invasive imaging can be performed on surfaces in their native states, and under near-physiological conditions. It has proven to be particularly successful for imaging biological samples such as proteins, nucleic acids and whole cells. By scanning, dynamic processes can be imaged such as erosion, hydration, physicochemical changes, and adsorption at interfaces. Therefore SFM is currently the SPM technique with the widest applicability for biomaterial research. In SFM a small tip attached to a micro-beam (cantilever) is scanned across the surface of a specimen, and deflected by topographic features (Fig. 8.8(a)). The force of interaction may be repulsive or attractive, giving rise to different modes of operation. Moving the cantilever from the interaction free zone far

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8.8 (a) Schematic drawing of the SFM principle, (b) Cantilever deflection versus sample-z-position-curve on a stiff surface monitored via SFM.

above the surface it snaps into contact (Fig. 8.8(b)) due to attractive force between the tip and sample, which can be described in a simple way by the Lennard Jones potential. The piezo pushes the tip further towards the sample and the positive repulsive force reaches a maximum. As the piezo is retracted the repulsive force is reduced, and the force changes sign. If the bending force of the cantilever becomes greater than the attractive force towards the surface, the tip loses contact. The tip can be held in the repulsive regimen of the Lennard Jones potential or oscillated in an attractive or repulsive regime resulting in different interactions.25,26 These differences are important, as biomolecules are deformed by applying a load of some nN as present in contact mode, i.e., in the repulsive regime.27 Deflection is usually monitored by a laser beam that is reflected and detected with a four split photodiode. This signal is used to maintain a constant force via a feedback loop, and to monitor the height data. SFM can be operated in a variety of modes that can provide different information about the sample. Usually the z-deflection is monitored, and interpreted according to the parameters under investigation. This is mainly done via monitoring of the zpiezo voltage. This can cause ambiguities, as piezo crystals exhibit hysteresis. This is overcome by monitoring the distance separately via inductive or fiberoptic sensors.28 For dynamic modes29 the application of an additional oscillation to the cantilever by a piezo crystal has to be performed. Magnetically driven cantilevers are also used for actuation.30 The quality of SFM data is essentially determined by the cantilever and the tip, influencing the resolution of topography and force measurements. Micromechanical properties of the cantilever, and the shape and chemical composition of the tip, which comes into direct

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contact with the sample, are essential. High aspect ratios and small tip radii are desirable for imaging steep slopes and deep crevices. Depending on the mode of operation, different parameters have to be optimized, which can be realized according to the methods described in the following. Silicon and silicon nitride cantilever fabrication based on photolithographic techniques are well established. Metal based (Ni) cantilevers31 and cantilevers made of piezoelectric material (lead zirconate titanate) are also produced for independent actuation and sensing.32 Cantilevers of various shapes (e.g. rectangular or V-shaped) and dimensions (usually 100 to 200 m in length) are available. New approaches to further miniaturization of the system have been made on microfabricated aluminum probes with length scales of 9 m.33±36 Coatings (Cr-Au, Pt, Al, TiN, W2C, TiO2, Co, Ni, Fe, Au with biological coating) can be applied to the cantilever to modify it against corrosion in the liquid phase, or for different applications when conductive, magnetic, or biological properties are necessary. The resonance frequency of the cantilever ranges from several kHz to several hundred kHz. Cantilevers with high resonance frequencies are used in dynamic mode as the tip oscillates at several hundred kHz above the surface. The stiffness of the cantilever is defined by the force constant k, and ranges between 0.01 to 100 N/m for the vertical deflection. Soft (k < 0.1 N/m) cantilevers are used in contact mode to minimize disturbance of the sample. Rigid cantilevers with a force constant larger than 1 N/m are used in non-contact or dynamic modes since they exhibit high resonant frequencies and small oscillation amplitudes of several nanometers. The force constant for lateral twisting can be determined for friction measurements,37 and the movement of the cantilever has been modeled accordingly with finite element analysis.38 The mechanical behavior and the determination of the spring constant is well described in the literature.39±46 The tip itself can consist of different materials, or is coated according to the application. For some applications, a hard surface is necessary and a diamondlike coating (DLC) is applied.31 The production of DLC coatings has been described,47 but other functionalities48±50 can be applied. Thus cantilever tips can be modified by chemical51 and biochemical52 functionalization. With silanization,53,54 or via thiols, self-assembled monolayers (SAMs) allow further modification.50 Proteins and bacteria55 can also be attached to the tip. Additional materials which vary the shape of the tip can be deposited, such as polystyrene, borosilicate and silica spheres, C60 molecules,56 carbon-nanotubes in general57±60 or single wall nanotubes with diameters of <3 nm61 and functionalized nanotubes.62 The first can be used to give a defined contact area. This is important because tips produced by lithographic methods are of pyramidal shape but do not present the necessary sharpness for high-resolution images. As the tip convolutes with surface features,63 its shape accordingly determines lateral resolution and the width with the possibility of entering small features. Electrochemical etching can produce sharper tips with an open angle

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near the apex of 20ë. For sharper tips an extra tip is grown on the apex of the conventional pyramidal tip using electron beam deposition (EBD)64 which leads to a curvature radius of 2 to 20 nm, or the above-mentioned nanotubefunctionalized tips are used. Real tip shape can be calculated by deconvolution of the experimental data.65±68 The resolution of steep slopes can be improved additionally by slow scanning and repeatedly z positioning at each pixel.69 Usually the typical scan range is up to 150 m but larger scan areas can also be reached with a scan area up to 25 mm2.69 In contact mode, topography and roughness are characterized by monitoring the vertical excursion of the cantilever in contact with the surface. Assigning a deflection signal in the z-direction to each point while scanning leads to the image. A typical example of a biological surface is shown in Fig. 8.9. Roughness can be calculated from the acquired data. Friction force microscopy is done by monitoring the twist of the cantilever while scanning over the surface.70 With this nanomechanical measurement, shear strength can be determined; an overview has been reported of friction on the atomic scale and is available.71 After chemical modification, the tip of the cantilever can be used to gain information about chemical properties, respectively through forces between a biomolecule and a biomaterial surface. Images are produced that display predictable contrast, and correspond to the spatial distribution of functional groups on the sample surface. Surfaces can be imaged in friction mode down to a resolution of about 0.5 nm.20 The different

8.9 Typical contact mode SFM image of bovine enamel after `etching' in HCl solution.

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chemical interactions of model surfaces with different chemical modifications is available.72 Image contrast is achieved between regions terminated by different functional groups,73 but this can be influenced by topographical features. Hydrophobic and hydrophilic parts of the surface have been distinguished74 and it has even been possible to gain information about chirality,75 which in the latter case was done on immobilized molecules by investigation of friction and adhesion forces. The load dependency is another important parameter. The atomic stick/slip mechanism, and interfacial friction, adhesion and elastic properties measured by friction force microscopy have been reviewed.76 In dynamic modes77±79 the cantilever is oscillated near its resonance frequency by an additional piezo crystal. The amplitude signal of the oscillating cantilever yields information about the interaction. A distinction between the non-contact and dynamic mode in the repulsive regimen is often made. In real non-contact mode, attractive forces are measured while the tip is oscillated above the sample with an amplitude of about 10 nm, never touching the sample surface. In the dynamic mode that is more commonly used, such as intermittent contact or tapping mode,80 the tip is oscillated with an amplitude of 20 to 100 nm and touches the surface at some point. The feedback loop works as in contact mode, maintaining constant amplitude. Topography and roughness information are obtained as described above. Additionally and simultaneously, phase image can be recorded, which reflects the phase shift between the generating and the response signal. If the cantilever oscillates freely in air, the waveform signal detected by the photodiode should be in phase with the signal driving the piezo element. This signal is sensitive to the interaction force between the cantilever and the surface. Phase contrast arises from the different amounts of energy dissipated by the tip while oscillating on surface areas with different elastic properties. A short overview on experiments and theoretical descriptions is available.81 Mechanical properties82±84 as well as chemical differences can yield a phase lag, which has been correlated quantitatively to the chemical interactions on chemically microstructured surfaces.81,85 In addition, SFM is also used for applications in which no scanning is required, where the tip is used as a sensitive force sensor (force spectroscopy).86±88 Loads of 1 to 1000 nN can be applied with tips that have tip radii of up to one micron leading to an investigation of the first 0.1 to 10 nm layers.89 Forces on a larger scale can also be measured by using the force apparatus designed by Israelachvili;90,91 this will cover larger areas but the limitation to mica or a modification of it is indicated. However, absolute force values can be determined with high accuracy because the interaction area is well known. With the setup of an SFM, surface and adhesion forces92 between two surfaces and elasticity, respectively hardness, by indentation of the substrate can be obtained. The data can be investigated by plotting the bending of the cantilever versus the load applied to the sample. By indenting the surface, it is

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possible to gain information about elasticity and the Young's modulus,27,93 as well as time dependent processes happening on polymers.89 By monitoring the forces between the tip and the surface, chemical patterns on a surface can be characterized.49 To monitor hardness by nanoindentation, usually blunter tips are in use to be able to relate to macroscopic hardness testing. The Berkovich tip with a nominal radius of 150 nm is used to monitor hardness, which is defined as the ratio of the maximum load to the projected contact area.94 By force mapping, two-dimensional arrays of force curves are recorded whilst scanning the tip across the sample. By this, the elastic properties can be monitored over the whole scan area. From the force curves, the Young's modulus can be calculated according to Sneddon's modification of the Hertzian model for elastic indentation.95

8.4

Optical microscopies

There is a huge variety of optical microscopies and spectroscopies which are all mainly used to study biological molecules, but not so often biomaterials. A comprehensive treatment of these methods is outwith the scope of this chapter. Therefore only three important techniques are outlined here in a short overview.96 Fluorescence microscopy is one of the most widely used techniques in biology. Biological molecules are labelled with a fluorescent marker (or label), i.e., are chemically modified, if they do not contain fluorescing entities like tryptophan themselves. A fluorescence microscope is a classical optical microscope in which different wavelengths can be detected visually, usually by selecting a wavelength by putting a filter between sample and detector. One advantage of fluorescence microscopy is the fact that some dyes quench the fluorescence of another dye by the so-called FoÈrster energy transfer. If two parts of a large biomolecule or two different molecules are labelled with these two dyes, one can study the distance dependence of the interaction. Another advantage is that the fluorescence wavelength is influenced by the environment; this can also be used to get more insight in the structural properties of a biomolecule. This method is, therefore, mainly used for studies on molecules attached to a biomaterial surface but not to study the biomaterials themselves. As outlined in section 8.3 the resolution is limited to that of an optical microscope, i.e., to some hundred nanometers. In confocal microscopy97 a laser beam is split and refocused on the plane of interest. This reduces one problem of normal optical microscopy where the entire sample is illuminated and in-focus and out-of-focus points contribute equally to the signal. In confocal microscopy, the sample is illuminated by a point source of light in the focus. The resolution reaches 200 nm in the xy-plane and 500 nm in the z-plane. The greatest advantage is the possibility of making three-dimensional maps of the sample to within a depth of around 100±200 m.

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There are three imaging modes in confocal microscopy: confocal epitransmission, where the backscattered light from within the specimen is imaged, confocal reflectance with light reflected from an opaque sample surface and confocal epifluorescence with light emitted by fluorescence. There are new developments (i.e., not in routine use) such as 4-microscopy which enhance resolution to around 40 nm.98 Scanning nearfield optical microscopy (SNOM) 99,100 is an optical microscope which is not operated in the far field as classical microscopies, but in the near field. Because the resolution limitations result from the diffraction of light in the far field, they are not valid in the near field. In the near field, the short-range interactions of an evanescent light field with a sample surface are detected. There are several practical solutions to near field microscopy. The first and most commonly used method uses a glass fiber (often metal coated) which is sharpened at its end so that light from a laser source at the other end of the fiber cannot propagate through the small exit hole and an evanescent field evolves. If this fiber is at a short distance from the sample surface, the evanescent near field interacts with the surface. The reflected or transmitted light is then collected in the far field. The resolution limit is given only by the geometrical parameter of the fiber opening. Resolutions at around 100 nm are commonly reached. The approach to the sample surface is often realized by wobbling the fiber and detecting the wobbling amplitude by a sensor such as a tuning fork. The amplitude is decreased when the fiber comes close to the surface. There are many new developments which aim at resolving the main limitations of SNOM: the small transmittance of the fiber has to be overcome and hence the very low intensities. This is particularly true if infra-red light is used, for which glass fibers are not transparent. Furthermore, the approach to the surface should be improved, e.g., by combining SNOM with SFM by channeling laser light through a hollow SFM tip instead of the optical fiber. The third limitation is the interaction radius which should be reduced to near-atomic dimensions. These limitations have so far hindered the common use of SNOM in the biomaterials community, but this will certainly change as soon as the technical solutions lead to more user-friendly instruments.

8.5

Future trends

There are several future trends to be foreseen. The first is to further increase the lateral resolution of techniques which can be used for in-situ studies, in particular in studies of the interface between biomaterials and biological materials, where optical techniques are superior to others. This has to be accompanied by practical solution of problems such as the need for flat surfaces for scanning probe methods, whereas biomaterials often exhibit rough surfaces. Furthermore, the development of techniques with very high time resolution in

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the femtosecond or even attosecond time regime in physics will extend more and more into biological research. This is already true for time-resolved spectroscopies, but not for time-resolved microscopies. The combination of surface analytical tools, together with biological knowledge and biochemical or microbiological methods will be of key importance for the whole field. Therefore, interdisciplinary research will be even more demanding in biomaterial research than in other fields of research.

8.6

Further information

There is a huge variety of review articles on biomaterial characterization, e.g., refs 1, 2, 101, which are often concentrated either on one type of microscopy or on one type of material, e.g., refs 102, 103. A non-comprehensive list is already given in the text. On the other hand there are text books and handbooks on all of the above-mentioned techniques, usually not covering the special needs of biomaterial research, but explaining the methods with their theoretical and/or practical background. The third source of information is the internet. In particular the websites of biomaterials organizations can be used as entrance sites for further information, see, e.g., www.biomat.net, www.biomaterials.net, www.bmn.com, ssb.biomaterials.ch, www.dgbm-news.de, www.esb-news.org, www.biomaterials.org, www.biomaterials.ca, www.biomaterials.org.au, www.surfaces.org, wwwsoc.nii.ac.jp/jsdmd/index-e.shtml.

8.7

References

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11. L. Reimer, Optical Sciences 80: Energy-Filtering Transmission Electron Microscopy, Springer, Berlin, (1995). 12. H. Bethge and J. Heydenreich, Elektronenmikroskopie in der FestkoÈrperphysik, Springer, Berlin, (1982). 13. M. v. Ardenne, Zeitschrift fuÈr Physik, 109, 553±572, (1938). 14. M. Knoll and R. Thiele, Zeitschrift fuÈr Physik, 113, 260, (1939). 15. G. Valdre, Fundamental Properties of Nanostructured Materials, National School of the Condensed Matter Group, Rimini, Italy, Sept. 20±25, 1993, 99±110, (1994). 16. G. Binnig, H. Rohrer, C. Gerber and E. Weibel, Physical Review Letters, 49, 57± 61, (1982). 17. G. Binnig, C.F. Quate and C. Gerber, Physical Review Letters, 56, 930±933, (1986). 18. D. R. Louder and B. A. Parkinson, Analytical Chemistry, 297±303, (1995). 19. S.O. Vansteenkiste, M.C. Davies, C.J. Roberts, S.J.B. Tendler and P.M. Williams, Progress in Surface Science, 57, 95±136, (1998). 20. R.M. Overney and E. Meyer, Nature, 359, 133±135, (1992). 21. S.M. Smith and J.L. Gilbert, Proceedings of the Electrochemical Society, 94 (15), 229±240, (1994). 22. H.G. Hansma, Proceedings of the National Academy of Sciences of the USA, 96, 14678±14680, (1999). 23. C. Galli, M. Collaud Coen, R. Hauert, V.L. Katanaev, M.P. Wymann, P. GroÈning and L. Schlapbach, Surface Science, 474, L180±L184, (2001). 24. D.T. Pierce and P.R. Unwin, Analytical Chemistry, 64, 1795±1804, (1992). 25. S. Paulo and R. GarcõÂa, Surface Science, 471, 71±79, (2001). 26. F. Dubourg and J.P. AimeÂ, Surface Science, 466, 137±143, (2000). 27. L. Weisenhorn, M. Khorsandi, S. Kasas, V. Gotzos and H.-J. Butt, Nanotechnology, 4, 106±113, (1993). 28. M. Heyde, H. Sturm and K. Rademann, Surface and Interface Analysis, 27, 291± 295, (1999). 29. M. Radmacher, R.W. Tillman and H. Gaub, Biophysical Journal, 64, 735±742, (1993). 30. W. Han, S.M. Lindsay and T. Jing, Applied Physics Letters, 69, 4111±4113, (1996). 31. T. Hantschel, S. Slesazeck, P. Niedermann, P. Eyben and W. Vandervorst, Microelectronic Engineering, 57±58, 749±754, (2001). 32. Y. Miyahara, T. Fujii, S. Watanabe, A. Tonoli, S. Carabelli, H. Yamada and H. Bleuler, Applied Surface Science, 140, 428±431, (1999). 33. T.E. SchaÈffer, M. Viani, D.A. Walters, B. Drake, E.K. Runge, J.P. Cleveland, M.A. Wendman and P.K. Hansma, SPIE, 3009, 48±52, (1997). 34. D.A. Walters, J.P. Cleveland, N.H. Thomson, P.K. Hansma, M.A. Wendman, G. Gurley and V. Elings, Review of Scientific Instruments, 67, 3583±3590, (1996). 35. D.A. Walters, M. Viani, G.T. Paloczi, T.E. SchaÈffer, J.P. Cleveland, M.A. Wendman, G. Gurley, V. Ellings and P.K. Hansma, SPIE, 3009, 43±47, (1998). 36. M.B. Viani, T.E. SchaÈffer, G.T. Paloczi, L.I. Pietrasanta, B.L. Smith, J.B. Thompson, M. Richter, M. Rief, H.E. Gaub, K.W. Plaxco, A.N. Cleland, H.G. Hansma and P.K. Hansma, Review of Scientific Instruments, 70, 4300±4303, (1999). 37. Feiler, P. Attard and I. Larson, Review of Scientific Instruments, 71, 2746±2750, (2000).

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9

Nanoindentation A B M A N N , The State University of New Jersey, USA

9.1

Introduction

Nanoindentation is now widely recognized as the preferred method for testing thin film and surface mechanical properties. Since surface and interfacial properties are of great importance in biology, and especially in biomaterials, there is a growing interest in the use of nanoindentation to study biological materials and systems. Current studies have tended to focus on hard tissues and biomaterials, but soft tissues and viscoelastic biomaterials are of growing interest. In the coming years the importance of nanoindentation testing in biology and biomaterials research is likely to show a rapid increase. This chapter describes the fundamentals of nanoindentation testing including instrumentation and data analysis. Also highlighted are the many pitfalls that may befall the unwary researcher. Unless the researcher is aware of these there is a real danger that the results obtained will say more about the analysis routines that are used than the material that has been tested. The chapter details the current results of nanoindentation testing on both soft and hard tissues and biomaterials, and also the future direction of nanoindentation studies on biological systems. The use of nanoindentation methods to study skeletal and dental tissues is highlighted. The nanomechanics of soft, living tissues and polymeric biomaterials are also described. Nanoindentation is now being applied to this problem with studies examining connective tissues and polymers using both static and dynamic methods. The difficulties in using these methods are discussed.

9.2

Instrumentation

The main instruments used to examine nanomechanical properties of surfaces and thin films are based on point-probes. These have developed from two historically different methodologies, namely scanning probe microscopy, specifically atomic force microscopy (AFM) (e.g. Binnig et al., 1986), and microindentation (e.g. Blau and Lawn 1986). The two converge at a length scale between 10±1000 nm, which is equivalent to a load range of Ns to mNs. Point-

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9.1 Operation of a commercial AFM. The deflection of the cantilever is a measure of the force acting on the tip. Deflections of the cantilever are monitored using the 4-quadrant detector.

probe mechanical tests in this range are often referred to as nanoindentation since the indents have nanometer dimensions. These nanomechanical tests are often used in conjunction with complementary techniques such as micro-Raman spectroscopy, electron microscopy (SEM and TEM) and high-energy diffraction studies. AFMs have many similarities to nanoindentation instruments. An AFM utilizes a piezoelectric stack to move either a probe tip or the sample with subnanometer precision in the lateral and vertical planes. The probe itself is a sharp tip mounted on the end of a cantilever beam. To measure mechanical properties with an AFM, the standard configuration is a hard probe tip (such as silicon nitride or diamond) mounted on a cantilever (see Fig. 9.1). The elastic deflection of the cantilever is monitored either directly or via a feedback mechanism to measure the forces acting on the probe. In general, the forces experienced by the probe tip split into attractive or repulsive forces. As the tip approaches the surface it experiences intermolecular forces that are attractive, although they can be repulsive under certain circumstances (Israelachvili, 1992). Once in contact with the surface the tip usually experiences a combination of attractive intermolecular forces and repulsive elastic forces. Two schools of thought exist regarding the role of attractive forces when the tip is in contact with the surface. The first is often referred to as the DMT or Bradley model. It holds that attractive forces only act outside the region of contact (Bradley, 1932; Derjaguin et al., 1975; Muller et al., 1980, 1983). The second theory, usually called the JKR model, assumes that all the forces experienced by the tip, whether attractive or repulsive, act in the region of contact (Johnson et al., 1971). Most real AFM contacts lie somewhere between these two theoretical extremes.

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The fundamental difference between AFM and nanoindentation is that during a nanoindentation experiment an external load is applied to the indenter tip. This load enables the tip to be pushed into the sample creating a nanoscale impression on the surface, otherwise referred to as a nanoindentation. Unlike conventional indentation or microindentation tests where the analysis uses optical imaging of the indentation impression, nanoindentation methods have been developed that continuously record the load, displacement, time and contact stiffness throughout the indentation process. Analysis of this data is then used to find the mechanical properties of the sample without using optical imaging. This type of continuously recording indentation testing was first applied in the former Soviet Union (Ternovskii et al., 1973; Bulychev et al., 1975, 1976; Shorshorov et al., 1982; Bulychev and Alekhin, 1987). Subsequently, it was applied at the nanoscale in the early 1980s (Pethica, 1982; Newey et al., 1982), hence giving rise to the field of nanoindentation testing. Typically nanoindentation instruments include a loading system that may be electrostatic, electromagnetic or mechanical, along with a displacement measuring system that may be capacitive or optical. Schematics of some commercial nanoindentation instruments are given in Fig. 9.2a, b and c. One major advantage of nanoindentation over conventional microindentation testing is the ability to measure the elastic modulus (E) as well as the hardness (H) of a sample. E is obtained from the contact stiffness (S) during unloading (see Fig. 9.3). S and E are related by the following equation that appears to be valid for all elastic contacts (Kendall and Tabor, 1971; Pharr et al., 1992): p 2 9:1 S ˆ p Er A  A is the contact area and Er is the reduced modulus of the tip and sample as given by: 1 …1 ÿ t2 † …1 ÿ s2 † ˆ ‡ Er Et Es

9:2

where Et,  t and Es,  s are the elastic modulus and Poissons ratio of the tip and sample, respectively. Load-displacement curves (Fig. 9.3) provide a lot of information, but they are only part of the story and additional information is usually helpful in explaining the observed results. The basic nanoindentation set-up can be adapted to obtain additional information about the deformation processes occurring, for example, in-situ measurements of acoustic emissions (Weihs et al., 1992; Bahr et al., 1997, 1999) and contact resistance (Mann et al., 2000, 2002) have been performed to identify fracture events and phase transformations, respectively. Environmental control has also been used to examine the effects of temperature and surface chemistry on the mechanical behavior of nanocontacts (Lucas and Oliver, 1999; Syed Asif et al., 2000; Asif and Pethica, 1997; Mann and Pethica,

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9.2 Schematic diagrams of several commercial nanoindentation devices. (a) is by MTS, Oakridge, Tennessee, (b) by Hysitron, Minneapolis, Minnesota and (c) by Micro Materials Limited, Wrexham, UK.

1996). In the future this is likely to be of great importance during in-vitro studies of biological materials under simulated physiological environments. Microscopy (optical, SEM, TEM or AFM) of the nanoindentation is often needed to ensure the correct interpretation of the mechanical data. Optical techniques will often reveal the presence of median or lateral cracks (Cook and Pharr, 1990). Electron microscopy (SEM/TEM) and AFM can be used to examine even the smallest nanoindentations to look for `pile-up' or `sink-in'. Pile-up is the accumulation of material at the side of the indenter during the indentation. This tends to increase the contact area between the tip and sample. Sink-in is when material is displaced into the sample during the indentation giving a reduction in

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9.3 Basic shape of a load-depth curve showing the contact stiffness, S, during unloading.

the contact area. The principal problem with nanoscale microscopy techniques (AFM, TEM and SEM) is the difficulty in finding the nanoindentations. It is usually necessary to make large, `marker' indentations in the vicinity of the nanoindentations to be examined in order to find them (Page et al., 1992). SEM has been used to find extrusions due to phase transformations (Pharr et al., 1991) and AFM is good for identifying pile-up and sink-in around nanoindents. TEM is useful for examining what has happened sub-surface, for instance, the indentation induced dislocations in a metal (Robertson and Fivel, 1999) or the phases present under a nanoindent in silicon (Mann et al., 2000). Although with TEM there is the added difficulty of sample preparation, and the associated risk of observing artifacts. Recently, there has been considerable interest in the use of focused ion beams to cut cross-sections through nanoindents (Bradby et al., 2000). When used in conjunction with SEM or TEM this provides an excellent means to see what has happened in the subsurface region. It should also be noted that some instruments are capable of performing a nanoindentation test and subsequently an AFM style scan of the indented region with the same probe tip. This makes it much easier to find the indentation impression after the nanoindentation test.

9.3

Data analysis

The data analysis routines are based on continuum mechanics models for ideal plastic and elastic materials deformed when an axisymmetric tip is pushed into them. Ideal elastic-plastic materials have a linear stress-strain curve until they reach their elastic limit, then they yield plastically at a yield stress, Y0, that remains constant during the ensuing deformation (Tabor, 1951). In a twodimensional indentation problem the yielding occurs because the Huber-Mises criterion (von Mises, 1913) or the Tresca criterion (Tresca, 1864) has been reached. Using these criteria it is found that during yielding the mean pressure, Pm, across the end of the tip is related to the yield stress by:

230

Surfaces and interfaces for biomaterials Pm ˆ KYo

9:3

where K ranges from 2.6 for the Tresca criterion to 3.0 for the Huber-Mises criterion. During nanoindentation testing the `nanohardness', H, is defined as the peak load, P, applied during the nanoindentation test, divided by the contact area, A, of the nanoindentation projected into the plane of the surface (see eqn 9.4). The definition of H is the same as the definition of Pm in eqn 9.3. Hence, if the test material behaves ideally eqns 9.3 and 9.4 can be used to relate H and Y0 (in reality few materials behave ideally so this relationship is often not valid): P 9:4 A Note that for very small nanoindentations particularly into soft, viscoelastic materials (e.g. membranes and biofilms) the effective load (P) can be influenced by intermolecular forces between the sample and the tip. This can give an adhesive contact as described by either the JKR model or DMT/Bradley model described earlier. For nanoindentation tests conducted in air the condensation of water vapor at the tip-sample interface may also affect the adhesive force acting. Since many biological materials are hydrophilic this capillary effect can impact the nanoindentation results. The elastic analysis of nanoindentation data is based on continuum elastic models first developed in the 19th century for a hemisphere pushed into an elastic half-space (Boussinesq, 1885; Hertz, 1882). Later these were developed by Love (1929, 1939) and Sneddon (1965) who found that for a variety of axisymmetric punches pushed into an elastic half-space the displacement () and the applied load (P) are related by: Hˆ

P ˆ m

9:5

where  and m are constants that depend on each punch's geometry. The significance of this equation will be seen later. Because of the problems associated with creating nanoscale axisymmetric tips, pyramidal indenter geometries have now become standard during nanoindentation testing. The most common geometries are the three-sided Berkovich pyramid and cube-corner (see Fig. 9.4). The Berkovich pyramid is based on the four-sided Vickers pyramid commonly used during microindentation testing. For both the Vickers and Berkovich pyramids the cross-sectional area of the pyramid's base, A, is related to the pyramid's height, D, by: A ˆ 24:5D2

9:6

The cube-corner geometry is sometimes used to make very small nanoindentations because it is much sharper than the standard Berkovich pyramid. This makes it easier to initiate plastic deformation at very light loads. There are, however, several issues that should be considered with the cube-

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9.4 Geometry of nanoindentation tips.

corner geometry; it can wear down quickly and become blunt, and the standard analysis routines that are discussed here are based on elastic models that assume the displacement into the surface is small in comparison to the tip radius. For the cube-corner geometry this may not be the case since the effective radius is very small. To obtain mechanical properties from the load/displacement curve an analysis based on eqns 9.4 and 9.5 is used. Most nanoindentation instruments include standard analysis software packages which provide values of hardness, H, and elastic modulus, E. These are based on the methodology developed by Oliver and Pharr (1992). A cautionary note should be given here regarding the `real' geometry of the contact and the geometry assumed in the analysis routines; these can differ significantly. Any difference between them will lead to errors in the calculated contact area and, hence, the values of E and H. Nanoindentation load-depth curves have a basic shape that depends on the elastic modulus, yield stress and viscoelastic properties of the test material (and the probe tip). The loading section of the curve approximates a parabola (Hainsworth et al., 1996), provided there is no time-dependency and no discontinuities are observed (as shown by Fig. 9.3). The load-depth curve during unloading is not parabolic, but in general follows a relationship which is similar in form to eqn 9.5, that is: P ˆ … ÿ i †m

9:7

where  is the total displacement and i is the intercept of the unloading curve with the displacement axis as shown in Fig. 9.5. Equation 9.7 assumes the unloading curve is exhibiting purely elastic behavior; viscoelasticity, phase transformations and fracture events can cause significant deviations from this situation. Further to this, eqn 9.7 is valid only for relatively shallow (low depth to width ratio) impressions and a surface that approximates a flat, elastic halfspace. For nanoindentations with a Berkovich pyramid this is generally the case, but may not be for the cube corner geometry described previously. For a wide range of materials Oliver and Pharr (1992) found typical values for m in eqn 9.7 to be around 1.5. By using this value for m eqn 9.7 is fitted to the unloading curve. Then applying eqns 9.1 and 9.4 gives E and H for the test material. Equation 9.1 relates the contact stiffness, S, during the initial part of the

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9.5 Nanoindentation data analysis. The first methods of data analysis used a flat cylindrical punch approximation (a), but this was later replaced with a paraboloid (b).

unloading curve (see Fig. 9.5) to the reduced elastic modulus (Er) and the contact area (A) at the peak load. Equation 9.4 gives the hardness (H) as the peak load divided by the contact area. For evaluating both E and H, knowledge of the contact area, A, at the peak load is vital. An incorrect contact area will give the wrong values for E and H. This is the most common source of problems in the analysis of nanoindentation data. The contact area is generally defined as a function that depends on the contact depth, c , thus the contact area becomes A ˆ Ac …c †. For a perfect Berkovich pyramid tip this would give the relationship between Ac and c described by eqn 9.6. For real tips, which are never perfect, the following expansion is used: Ac …c † ˆ 24:5c2 ‡

7 X

Cj

p  2j c

9:8

jˆ1

where Cj are calibration constants of the tip. The contact depth, c , is not the same as the indentation depth because the surface around the indentation will be elastically deflected during loading. Sneddon's analysis (1965) for axisymmetric contacts allows for the deflection of the surface at the contact's edge. Subtracting the deflection from the total indentation depth at peak load gives the contact depth. For a paraboloid indenter geometry the elastic deflection at the edge of the contact, s , is given by: P P ˆ 0:75 9:9 S S where S is the contact stiffness and P the peak load. The paraboloid geometry was found by Oliver and Pharr (1992) to give the best results for nanoindentations using Berkovich tips. Using eqn 9.9 the contact depth at peak load becomes: s ˆ 

c ˆ  ÿ s

9:10

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Using the load/displacement data from the unloading curve and eqns 9.1, 9.2, 9.4, 9.7, 9.8, 9.9 and 9.10, the hardness and reduced elastic modulus for the test sample can be calculated. To find the elastic modulus of the sample, Es, it is also necessary to know Poisson's ratio, s , for the sample as well as the elastic modulus, Et, and Poisson's ratio, t , of the indenter tip. For a diamond tip these are 1141 GPa and 0.07, respectively. Finding the coefficients (Cj) in eqn 9.8 is vital to the analysis. SEM, TEM or AFM can be used, but they are time consuming and prone to errors. Oliver and Pharr (1992) suggested calibrating with a standard specimen that is mechanically isotropic and has a known E and H that does not vary with indentation depth (fused silica is the most popular standard, E ˆ 72GPa,  ˆ 0:17). To calibrate the tip geometry a series of nanoindentation tests to various depths are performed on the standard sample. From the unloading curve of each test the contact stiffness, S, at the peak load, P, and the contact depth, c , are found. Then provided E is known a priori, eqn 9.1 can be used to calculate the contact area, A, at the contact depth, c . By repeating this for many contact depths the tip area function, Ac(c ), is found. The machine stiffness (or compliance) must also be calibrated since the machine frame is also subjected to a mechanical load during a nanoindentation test. To calibrate the machine frame large nanoindentations are made in a soft material such as aluminum with a known, isotropic elastic modulus. For very deep nanoindentations made with a Berkovich pyramid the contact area, Ac(c ), can be reasonably approximated to 24.5c2 . Using this value of A eqn 9.1 is used to find the expected contact stiffness, S, for the material. Any difference between the expected value of S and the value measured from the unloading curve will be due to the compliance of the machine frame. Some of the limitations of the analysis routines have already been discussed but other problems arise because the analysis was developed to look at materials with similar properties to the calibration material (typically fused silica). If the test material is significantly different from the calibration material the analysis will be incorrect. This may be due to the sample material's mechanical anisotropy, time-dependent effects, residual stresses at the sample surface or changes in the nanoindentation's shape after elastic recovery. Residual stresses change the contact area by causing pile-up (compressive stress) or sink-in (tensile stress) around the nanoindentation (Tsui et al., 1996; Bolshakov et al., 1996), hence, eqns 9.9 and 9.10 are not valid. Similar contact area errors are often seen for thin films on a substrate (Tsui et al., 1999a, 1999b). A further complication is the elastic recovery of the deformed region during unloading. This means the geometry of the residual indentation impression in the material is not as expected by the Oliver and Pharr analysis. Once again this is due to differences between the mechanical properties of the test sample and the calibration material. The exact shape of an unloaded nanoindentation on a material exhibiting elastic recovery is not simply an impression of the tip shape,

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rather, there is some elastic recovery of the nanoindentation sides giving them a slightly convex shape. To deal with the variations in the recovered nanoindentation shape, it has been suggested (Hay et al., 1999) that during the analysis a tip geometry that has slightly concave sides be used. This requires a modification to eqn 9.1: r A 9:11 S ˆ 2Er  where is a correction term dependent on the tip geometry (Hay et al., 1999). The standard method for analyzing nanoindentation data is based on the unloading curve, virtually ignoring the loading curve data. The loading curve has been ignored because it depends on both elastic and plastic properties, while the unloading curve depends mainly on elastic properties. There are, however, some alternative analysis methods that consider the loading section of the loaddepth curve. For homogeneous samples it has been found during loading that P is linearly proportional to 2 (Hainsworth et al., 1996; McGurk and Page, 1999) with the constant of linearity depending on E and H. Analysis of the loading curve has yet to gain popularity, but this may change in the future. Another approach treats the nanoindentation curve as a plot of force against distance. Integration under the unloading curve gives the elastic strain energy while integration under the loading curve gives the total work of indentation. Subtraction of the two gives the energy lost in plastic deformation. This approach has been used on a material that work-hardens to estimate H/Er (Cheng and Cheng, 1998), but again this analysis has yet to be widely used. One significant improvement to the standard method of performing and analyzing nanoindentation data is adding a small AC load on top of the DC load (Pethica and Oliver, 1989; Oliver and Pethica, 1989). The AC load creates a dynamic system where the sample acts as a spring with a stiffness S (the contact stiffness), and the nanoindentation system acts as a series of springs and dampers. An analysis of the dynamic system gives: Pos q ÿ1 ÿ1 2 2 2 2 9:12 …!† ˆ f…S ‡ Cf † ‡ Ks ÿ m! g ‡ ! D tan…† ˆ

!D …S ÿ1 ‡ Cf †ÿ1 ‡ Ks ÿ m!2

9:13

where Cf is the load frame compliance (the reciprocal of the load frame stiffness), Ks is the stiffness of the probe tip support springs (typically in the region of 50±100 N/m), D is the instrument damping coefficient, Pos is the magnitude of the load oscillation, …!† is the magnitude of the displacement oscillation, ! is the oscillation frequency (typically 50±100 Hz), m is the mass of the indenter and  is the phase angle between the AC force and the displacement oscillation.

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In order to find S using either eqn 9.12 or 9.13, it is necessary to calibrate the dynamic response of the system when the tip is not in contact with a sample …S ÿ1 ˆ 0†. This calibration combined with the standard DC calibrations will provide the values for all of the constants in the two equations. …!† and  are both measured by the lock-in amplifier used to generate the AC signal. Since the S obtained is the same as the S in eqn 9.11 it follows that the Oliver and Pharr analysis can be applied to obtain Er and H throughout the entire nanoindentation cycle. The dynamic analysis as detailed here was developed for the NanoindenterTM, MTS, Oakridge, Tennessee, but a similar analysis has been applied to other commercial instruments, such as the TriboscopeTM, Hysitron, Minneapolis, Minnesota (Asif et al., 1999). If the test material is viscoelastic (as is the case for many biological materials) the stiffness, S ˆ S 0 ‡ S 00 , and the elastic modulus, E ˆ E0 ‡ E00 , are complex. The storage modulus, E0 , and loss modulus, E00 , can then be found by using eqn 9.1 to relate S and E. This method for measuring viscoelastic properties using nanoindentation has now been proven in principle on polymeric materials and will be very important in future studies of biological materials. The analysis methods detailed here were developed for isotropic materials where the elastic modulus is assumed to be either independent of direction or a polycrystalline average of a material's elastic constants. Many crystalline materials exhibit considerable anisotropy in their elastic constants, hence, the standard analysis techniques may not always be appropriate. The problem of a rigid indenter pressed into an elastic, anisotropic half-space has been considered by Vlassak and Nix (1993, 1994). They have shown that anisotropy in crystalline metals can change the measured elastic modulus by a factor of 2 and hardness by up to 20% between different orientations. Mechanical anisotropy is a particular issue in hard tissues and biomaterials because the bone mineral hydroxyapatite is very elastically anisotropic.

9.4

Thin films and coatings

In many biomaterials applications the surfaces are coated with thin films that are intended to increase wear resistance, biocompatibility or some other property. Measuring the mechanical properties of these films is one of the major applications of nanoindentation. There is an often quoted 10% rule. This states that nanoindentations in a film must be to a depth of less than 10% of the film's thickness if only the film properties are to be measured. This has no real validity (Page and Hainsworth, 1993). There are film/substrate combinations for which 10% is very conservative, while for other combinations even 5% may be too deep. When testing thin films the maximum depth of indentation relative to film thickness that can be used depends on the zones of the elastic and plastic strain fields. In general these should be contained in the film and not penetrate into the substrate if film-only properties are to be measured. Experimental results (Oliver

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et al., 1987) indicate that the effect of the substrate on the measured elastic modulus of the film can be quite different to the effect on hardness, this is due to the zones of the elastic and plastic strain fields being different sizes. In recent theoretical and experimental studies the importance of material pileup and sink-in has been extensively investigated. For thin film materials using the Oliver and Pharr (1992) method of analysis fails to account for the pile-up or sink-in and results in a large error in the values for E and H. There are three clearly identifiable factors affecting the pile-up and sink-in around nanoindentations during testing of thin films: · residual stresses · degree of work hardening · ratio of film and substrate mechanical properties. Many unprepared researchers have misguidedly taken the values of E and H obtained during nanoindentation testing to be absolute values, only to find out later that the values contain significant errors due to pile-up or sink-in affecting the contact area, A. It is always worth obtaining an AFM or SEM image of a few of the nanoindentations in order to assess if pile-up or sink-in is a problem. One method of analysis proposed by Joslin and Oliver (1990) looks at the ratio of E2 to H. This has the advantage of being independent of the contact area, A, and hence it should not be affected by pile-up or sink-in. Relatively little work has been done on measuring the mechanical properties of soft, viscoelastic films using nanoindentation methods. One of the few examples is a comparative study of mechanical properties by Payne et al. (1998) that used a dynamic nanoindentation method and a dynamic mechanical analyzer. Good agreement was found between the two methods provided the coatings were not too thin or brittle.

9.5

Hard biological materials

In recent years there has been a steady growth in the number of researchers using advanced mechanical characterization techniques to study bone (Rho et al., 1997, 1999). Some of the results are summarized in Table 9.1. For comparison, data obtained by the author's research group on animal bone are included in Table 9.1. Nanoindentation testing of bone in a liquid environment (Rho and Pharr, 1999; Hengsberger et al., 2002) has been used to simulate physiological conditions. This is done using a liquid cell with a small electric heater. Care must be taken to ensure that the liquid is wetting the indenter shank when nanoindentation tests are performed in liquids. When the liquid is wetting the shank (contact angle  0) the liquid meniscus introduces an extra downwards force that is constant and can be compensated for by the instrument. A non-zero contact angle will give a varying capillary force and reliable testing is impossible (Mann and Pethica, 1996). When performing dynamic testing in a liquid the analysis given earlier must be

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Table 9.1 Nanoindentation date for bone from humans and a variety of mammals. The results show the effects of hydration on bone mechanical properties and also the variation between different species. The results for mouse bones indicate that the presence/absence of the bone protein osteopontin has little effect on the mechanical properties of young bones (3 week old) Study

Bone type

Rho et al. (1999)

Dry vertebrae and tibiae interstitial lamellae (longitudinal direction) Dry osteons (longitudinal direction) Dry trabeculae (longitudinal direction) Dry osteon and interstitial lamellae average (transverse direction) Dry trabeculae (transverse direction)

Hengsberger et al. (2002)

Dry trabecular bone Dry compact bone Wet trabecular bone at 37 ëC Wet compact bone at 37 ëC

Mann et al. Dry sheep compact bone (unpublished) Hydrated sheep compact bone (room temperature) Dry sheep cancellous bone Dry cow compact bone Dry cow cancellous bone Mann et al. Wildtype compact mouse bone (unpublished) (3 week old) Osteopontin deficient mouse bone (3 week old)

Elastic modulus (GPa)

Hardness (GPa)

25.7  1.7 22.4  1.2 19.4  2.3 16.6  1.1

Range for all data 0.52±0.74

15.0  2.5 21 to 27 17 to 23 8 to 20 9 to 20

Not available

18.89 16.20

0.814 0.59

13.35 25.54 14.53

0.53 1.09 0.59

38.42  6.84

2.12  0.57

37.89  7.18

2.08  0.57

modified further as the viscosity of the liquid will introduce an additional damping term. In this situatuion the instrument damping, D in eqns 9.12 and 9.13, must include the damping due to the liquid. This can be done during the calibration of the instrument prior to testing. The calibration must be performed at the same frequency and oscillation amplitude as that used during the subsequent testing and the indenter shaft must be immersed to the same depth as that used during the liquid testing. There is also a small change in the indenter's mass due to buoyancy in the liquid, however, a simple calculation reveals that this is negligible in all practical circumstances. The differences in mechanical behavior between compact or cortical bone and trabecular or cancellous bone have been the subject of several investigations (Turner et al., 1999; Zysset et al., 1999). The compact bone is generally found to

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have higher elastic modulus and hardness than the trabecular bone; in both cases the mechanical properties are reduced by hydration (Rho and Pharr 1999; Hengsberger et al., 2002). There is also a clear difference between the mechanical properties of osteonic lamellae and interstitial lamellae. The interstitial lamellae have a higher elastic modulus because of their greater degree of mineralization when compared to osteonic lamellae (Dorlot et al., 1986). Other studies have looked at elastic anisotropy in bone. These have shown a substantial variation in elastic modulus with direction of nanoindentation. For osteonic lamellae there is a variation with orientation from a minimum of 14.9  2.0 GPa to a maximum of 25.1  2.1 GPa, while for interstitial lamellae the range for different orientations is 18.5  1.1 GPa to 27.1  1.7 GPa (Fan et al., 2002). Significantly there appears to be little change in the nanoscale mechanical properties of human bone with age of the individual (Hoffler et al., 2000; Rho et al., 2002). This suggests that the loss of bone mass with age does not result in a change in the local mechanical properties of the bone. A similar result is observed in rat osteopetrotic mutations (Jamsa et al., 2002) and when using TEM to study the nanostructure of osteoporitic bone it is found to be very similar to normal bone (Rubin et al., 2003). These results are significant as they suggest that many conditions that change the density of bone do not affect its local mechanical behavior, however, these studies have not examined the bone's viscoelastic properties. It is possible that the mineral phase in osteoporitic bone is the same as in healthy bone, but that the collagen content or bonding of the collagen and other organic components is different, this would have little affect on E or H, but might change the viscoelastic behavior. There have been several studies of dental enamel (Xu et al., 1998; Cuy et al., 2002) using depth-sensing indentation tests. These have shown that dental enamel is elastically and plastically anisotropic, as would be expected for a material that is largely hydroxyapatite, which is itself elastically anisotropic. There are also substantial variations in the enamel's mechanical properties between the occlusal surface and the enamel-dentin junction (see Table 9.2). The enamel-dentin junction is of considerable interest (Marshall et al., 2001a) since dentin is generally more compliant than enamel. A reduction in the enamel's elastic modulus close to the enamel-dentin junction is necessary to avoid a discontinuity in elastic properties at the interface. A discontinuity in elastic properties would increase the risk of the interface failing. Dentin and bone, unlike enamel which is 95% inorganic or mineral (Williams et al., 1989), contain a significant quantity of organic material including collagen and are porous. Dentin for instance contains tubules that are filled with fluid. Nanoindentation of dentin has found large local changes in mechanical properties, particularly close to dentinal tubules (Kinney et al., 1996) where the walls of the tubule show an increase in the degree of mineralization. Dentin and bone are also `living' and can over time respond to changes in their mechanical loading. In contrast, once a tooth has erupted the ameloblasts that produce enamel are lost.

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Table 9.2 Nanoindentation data for an upper second molar (Cuy et al., 2002). There is a large change in E and H between the occlusal surface and the enamel-dentin junction (EDJ) Position Lingual occlusal surface Lingual EDJ Center occlusal surface Center EDJ Buccal occlusal surface Buccal EDJ

Elastic modulus (GPa)

Hardness (GPa)

108 52 112 85 100 55

5.5 2.5 5.9 4.1 5.2 2.6

Apart from chemical and physical interactions in the mouth, the enamel will remain unchanged for the rest of its natural life. An acidic oral environment, for instance due to the presence of acidogenic bacteria, can lead to the development of carious lesions in the dental enamel. These are the precursors to caries or cavities. Nanoindentation results across the depth of a carious lesion show how the hardness and elastic modulus varies with position from the surface (Fig. 9.6). The variations are due to changes in the degree of mineralization. The interior of the lesion is very demineralized and as a result it has a low hardness and elastic modulus. Close to the surface the lesion is only partially demineralized and as a result significantly stronger than the lesion's interior. This is in good agreement with the cariology of lesions as described by Newbrun (1989).

9.6 Mechanical properties across a carious white spot lesion (the mean position is the distance below the enamel's surface, so 0 is at the enamel's surface).

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For all mineralized tissues it has been found that there is a significant effect on their mechanical properties due to the environment they are tested in. For bone, tests in aqueous and simulated physiological conditions can change the hardness and elastic modulus by 20% (Hengsberger et al., 2002; Rho and Pharr, 1999). For enamel and dentin the difference between the `dry' and `wet' mechanical properties can be 10% (Balooch et al., 1998; Marshall et al., 2001b). For all mineralized tissues there is a risk of demineralization when they are immersed in a liquid such as distilled water, so when testing in liquid care must be taken in choosing an appropriate solution.

9.6

Soft biological materials

Nanoscale mechanical characterization of soft tissues and biomaterials is likely to be an area of future interest, but at the present time very little work has been done in this area. Some of the rare examples of this type of characterization are studies on cartilage repair tissue (Ebenstein et al., 2004) and physisorbed salivary pellicle (Dickinson and Mann, 2004). These materials exhibit distinctly different indentation curves from those described earlier. This is principally due to time-dependent effects and adhesion between the indenter tip and the sample, a schematic example curve is shown in Fig. 9.7. Due to significant differences between the load-depth curves of Figs 9.5 and 9.7 it is not appropriate to use the Oliver and Pharr analysis for viscoelastic, adhesive materials. The dynamic analysis outlined earlier can be used to obtain the storage and loss modulus. Alternatively, some properties can be obtained from the load-depth curve such as unloading stiffness, resistance to penetration and volumetric creep (Ebenstein et al., 2004). It should be noted that these properties are not materials properties in the sense of elastic modulus and hardness, but they do provide a measure of the material's behavior under mechanical loading.

9.7 Schematic nanoindentation curve for a viscoelastic, adhesive biomaterial.

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9.8 Nanoindentation of salivary pellicle showing significant creep and adhesion as would be expected for a hydrophilic, viscoelastic protein film. (a) shows normal pellicle while (b) shows the effect of exposure to tea (tannins).

Thin, physisorbed protein films are of great importance in biomaterials. They may appear during the early stages of infection, encapsulation or integration of an implant. Deposition of protein films is also seen in biofouling of sensors and occurs naturally in the mouth were salivary proteins form a pellicle layer on the surface of dental enamel. Figure 9.8 shows nanoindentation curves for salivary pellicle deposited in-vivo on enamel over a period of two hours.

9.7

Conclusion

Nanoindentation is now the standard technique for characterizing the mechanical properties of thin films and surface, as such it will become of growing importance to researchers interested in biological materials. The methods were originally designed for the characterization of relatively hard and rigid materials like metals and ceramics. Recently the utility of the methods has been extended to softer, viscoelastic materials such as polymers. This has required a re-assessment of the traditional methods of data analysis and has directly led to the development of dynamic analysis methods to measure storage and loss modulus. These dynamic methods have been successfully applied to some polymers, but they have yet to find widespread use. For the study of many biological tissues the use of dynamic analysis will be vital. There is a further complication for biological materials which relates to the environment in which they are tested. The human body is essentially an aqueous system at 37 ëC; for realistic in-vitro nanoindentation testing similar environmental conditions must be used. This has been achieved by performing nanoindentation testing in a heated liquid environment, but this introduces some instrumental complications related to thermal drift and evaporation. Care must be taken if these problems are to be avoided. As an example, a sealed chamber

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will reduce evaporation and the first few indentations in a large run should be ignored as they are those most prone to thermal drift. For some hydrated biomaterials an aqueous environment is not needed, but the environmental humidity should still be monitored as the sample may become de-hydrated if the humidity is low. There is the further issue of adhesion as shown by Figs 9.7 and 9.8. Adhesion between the tip and sample is common in hydrophilic, viscoelastic biological materials. One possible way to avoid this problem is to hydrophobe the tip with surfactants. This will reduce the adhesion, but over the course of many indentations the surfactant layer may be fully or partially removed, hence changing the perceived mechanical properties. One area in need of further improvement is the data analysis routines. These would ideally incorporate viscoelasticity, adhesion and the effects of pile-up and sink-in. To achieve this it is likely that some theoretical advances in contact mechanics modeling will be needed. This may ultimately require the incorporation of computer simulations into the analysis routines. The combination of nanoindentation and in-situ scanned imaging has been extremely helpful in addressing the errors associated with an incorrect contact area. However, the whole concept of contact area is confused by the combined effects of viscoelasticity and adhesion which cause the contact area to be continuously changing during the nanoindentation.

9.8

Further information

Nanoindentation has advanced rapidly over the past two decades and there are now a number of articles (Haque, 2003), book chapters (Mann, 2004) and books (Fischer-Cripps, 2002) that describe the technique in detail. There have also been two focus issues of the Journal of Materials Research (JMR) dedicated to nanoindentation techniques (June 1999, Vol. 14, Issue 6 and January 2004, Vol. 19, Issue 1) and an ongoing series of Materials Research Society Symposia (`Fundamentals of Nanoindentation and Nanotribology I-III'). The proceedings of these symposia and the focus issues of JMR are particularly useful sources of information for the latest developments in nanoindentation research. Two articles by Oliver and Pharr (1992, 2004) warrant specific mention as they discuss the methods used to analyze nanoindentation data and highlight the recent developments in this area.

9.9

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Rho, J.Y. and Pharr, G.M. (1999), `Effects of drying on the mechanical properties of bovine femur measured by nanoindentation', Journal of Materials Science-Materials in Medicine, 10, 485±488. Rho JY, Tsui TY and Pharr GM (1997), `Elastic properties of human cortical and trabecular lamellar bone measured by nanoindentation', Biomaterials, 18, 1325±1330. Rho, J.Y., Roy, M.E., Tsui, T.Y. and Pharr, G.M. (1999), `Properties of microstructural components of human bone tissue as measured by nanoindentation', Journal of Biomedical Materials Research, 45, 48±54. Rho, J.Y., Zioupos, P., Currey, J.D, and Pharr, G.M. (2002), `Microstructural elasticity and regional heterogeneity in human femoral bone of various ages examined by nano-indentation', Journal of Biomechanics, 35, 189±198. Robertson, C.F. and Fivel, M.C. (1999), `The study of submicron indent-induced plastic deformation', J. Mater. Res., 14, 2251±2258. Rubin, M.A., Jasuik, I., Taylor, J., Rubin, J., Ganey, T. and Apkarian, R.P. (2003), `TEM analysis of the nanostructure of normal and osteoporitic human trabecular bone', Bone, 33, 270±282. Shorshorov, M.Kh., Bulychev, S.I. and Alekhin, V.P. (1982), Sov. Phys. Doklady, 26, 769. Sneddon, I.N. (1965), `The relationship between load and penetration in the axisymmetric Boussinesq problem for a punch of arbitrary profile', Int. J. Eng. Sci., 3, 47±57. Syed Asif, S.A., Colton, R.J. and Wahl, K.J. (2000), `Nanoscale Surface Mechanical Property Measurements: Force Modulation Techniques Applied to Nanoindentation', in Interfacial Properties on the Submicron Scale, J. Frommer and R. Overney, eds, ACS Books. Tabor, D. (1951), Hardness of Metals, Oxford University Press, Oxford. Ternovskii, A.P., Alekhin, V.P., Shorshorov, M.Kh., Khrushchov, M.M. and Skvortsov, V.N. (1973), Zavod Lab., 39, 1242. Tresca, H. (1864), Cr. Hebd. Acad. Sci., 59, 754. Tsui, T.Y., Oliver, W.C. and Pharr, G.M. (1996), `Influences of stress on the measurement of mechanical properties using nanoindentation. 1. Experimental studies in an aluminum alloy', J. Mater. Res., 11, 752±759. Tsui, T.Y., Vlassak, J. and Nix, W.D. (1999a), `Indentation plastic displacement field: Part I. The case of soft films on hard substrates', J. Mater. Res., 14, 2196±2203. Tsui, T.Y., Vlassak, J. and Nix, W.D. (1999b). `Indentation plastic displacement field: Part II. The case of hard films on soft substrates', J. Mater. Res., 14, 2204±2209. Turner, C.H., Rho, J., Takano, Y., Tsui, T.Y. and Pharr, G.M. (1999), `The elastic properties of trabecular and cortical bone tissues are similar: results from two microscopic measurement techniques', Journal of Biomechanics, 32, 437±441. Vlassak, J.J. and Nix, W.D. (1993), `Indentation modulus of elastically anisotropic halfspaces', Phil. Mag. A, 67, 1045±1056. Vlassak, J.J. and Nix, W.D. (1994), `Measuring the elastic properties of anisotropic materials by means of indentation experiments', J. Mech. Phys. Solids, 42, 1223±1245. von Mises, R. (1913), Gottinger Nachr. Math Phys Klasse, 582. Weihs, T.P., Lawrence, C.W., Derby B., Scruby, C.B. and Pethica, J.B. (1992), `Acoustic emissions during indentation tests', MRS Symp. Proc., 239, 361±366. Williams, P.L., Warwick, R., Dyson, M. and Bannister, L.H. (eds) (1989), `Splanchnology: The teeth', in Gray's Anatomy, Churchill Livingstone, New York, pp. 1308±1309.

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Xu, H.H.K., Smith, D.T., Jahanmir, S., Romberg, E., Kelly, J.R., Thompson, V.P. and Rekow, E.D. (1998), `Indentation damage and mechanical properties of human enamel and dentin', Journal of Dental Research, 77, 472±480. Zysset, P.K., Guo, X.E., Hoffler, C.E., Moore, K.E. and Goldstein, S.A. (1999), `Elastic modulus and hardness of cortical and trabecular bone lamellae measured by nanoindentation in the human femur', Journal of Biomechanics, 32, 1005±1012.

10

Surface plasmon resonance V H P EÂ R E Z - L U N A , Illinois Institute of Technology, USA

10.1 Introduction Surface plasmon resonance spectroscopy (SPR) has emerged, over the last two decades, as a powerful surface analytical technique for the study of biomolecular and biomolecule-surface interactions. It was probably the demonstration of its sensitivity for protein adsorption studies (Kooyman, 1988; Stenberg, 1991), the facile construction of SPR instruments, and the introduction of commercial SPR equipment (JoÈnsson U., 1991; MeleÂndez, 1996 and 1997; Rich, 2001) that spurred such interest in this technique over a broad range of disciplines and research groups around the world. One of the earliest reports on the SPR phenomenon dates back to the beginnings of the 20th century when dark and light bands were observed in the background when a continuous source irradiated a metal grating with light polarized in the plane of incidence (Wood, 1912). This was subsequently explained as the excitation of surface electromagnetic waves at the metal-air interface (Fano, 1941). Years later, the concept of volume plasmons in metals was introduced. These volume plasmons were described as longitudinal fluctuations of volume electron density. The concept of fluctuating electron density allowed the description of volume plasmons in a theoretical format (Ritchie, 1957). In 1959, an experimental confirmation of Ritchie's theory was reported (Powell, 1959). Ritchie's theory also described `lowered' plasmon modes at metallic foils boundaries. These `lowered' plasmon modes were later described as surface plasmons (Stern, 1960). It was not until the early 1980s that the first demonstrations on the use of SPR for biosensing applications were first reported (Leidberg, 1983; Flanagan, 1984) and it took almost another decade for the advent of commercial SPR instruments (BIAcore, IASYS, IBIS, SPREETA, Bio Tul, GWC Instruments, and others). Today, research laboratories around the world routinely use SPR to study adsorption, biospecific molecular recognition, self-assembly, and other phenomena occurring at the surface of metallic thin films. The SPR phenomenon is also responsible for the unique optical properties of colloidal metals such as gold. It was Michael Faraday who first proposed,

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through a series of qualitative experimental observations, that the bright red color of ruby glass and his colloidal gold preparations was due to finely divided gold particles (Faraday, 1857). The optical properties of colloidal gold (and other metals) have only recently begun to be exploited for colorimetric detection assays (Elghanian, 1997; Storhoff, 1998; Kim, 2001; Otsuka, 2001; Berchmans, 2002; Thanh, 2002) and the analysis of biospecific interactions (Nath, 2002). It is likely that novel applications of metallic nanoparticles exploiting the SPR phenomenon will emerge in the near future. Thus, offering complementary advantages to planar substrates and/or an entirely new set of nanomaterials with unique optical properties (PeÂrez-Luna, 2004). Surface plasmon resonance finds widespread uses in biochemistry, biomaterials and surface science. However, a detailed understanding of this technique requires knowledge of optics, solid-state physics, and surface chemistry. Often, the user has only a superficial understanding of these fields. This chapter presents the basic principles of the SPR phenomenon, the tailoring of surfaces to immobilize receptors, and a representative description of applications that this technique has encountered in research laboratories. The use of thin metallic films is emphasized because it is the component of commercial instruments. However, the recent interest in the use of metallic nanoparticles for bioassays warrants a brief description of these systems. Thus, their optical properties and applications are also discussed in this chapter.

10.2 Surface plasmon resonance phenomenon The free electrons responsible for the conductive properties of metals move freely and continually throughout the material. The collective excitation of these electrons gives rise to fluctuating or oscillating electron density regions that receive the name of plasmons and occur with characteristic frequencies. Thus, such oscillations occur with quantized energy of frequency given by (Feldman, 1986) s 4ne e2 P ˆ 10:1 ! mo Here !P is the frequency of vibration of the electrons in the plasma wave, and ne , e and m0 are the electron density, charge and mass of the electron respectively. At the surface of the metal the kinetic energy of the electrons can cause them to move away from the bulk. However, when the electrons separate from the bulk of the metal their potential energy increases. This increase in potential energy draws them back to the bulk of the metal so that they oscillate perpendicular to the surface of the metal while traveling along the surface. This is analogous to the waves in the ocean where water travels along its surface but oscillates perpendicular to it. At the surface of the metal, the dielectric medium

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in contact with the metal modifies the frequency of vibration of surface plasmons such that their frequency becomes !P SP ˆ p ! 1 ‡ d

10:2

SP is the frequency of vibration of surface plasmons and d is the where ! dielectric constant of the medium adjacent to the metal film. Surface plasmons can be excited with electromagnetic radiation provided that the momentum and frequency of the photons matches those of the plasmons. Therefore, in order to excite the surface plasmons it is necessary for the electric component of light to oscillate in the plane of incidence (perpendicular to the surface of the metal). This state of polarization of light is called p-polarization and is required in order to excite the surface plasmons of metallic thin films. The wave vector for surface plasmons at a metal-dielectric interface is given by r 1 1  ˆ ckSP ‡ 10:3 ! m d In this equation, the dielectric constant of the metal, m , is a strong function of frequency. An adequate representation of the free electron behavior of metals is represented by the Drude model (Kittel, 1956), which relates the dielectric constant of the metal, m , to the frequency as follows m ˆ 1 ÿ

!2P !  ‡ i! d † !…

10:4

P represents the bulk Here, 1 is called the high frequency dielectric constant, ! d ˆ vf =Rbulk is the relaxation or damping frequency, plasma frequency, and ! where vf is the velocity of electrons at the Fermi level and Rbulk the mean free path of conduction electrons. The dielectric constant is thus a complex number where the imaginary part accounts for attenuation of electromagnetic radiation in metals.

10.2.1 Thin films The surface plasmons on the surface of metals can be excited by electromagnetic radiation provided that the incident light is p-polarized and the momentum and frequency of the plasmons matches those of the incident photons. Figure 10.1, shows a plot of the wave vector for surface plasmons, ksp, versus frequency, ! (solid line). This curve was obtained from combining eqns 10.3 and 10.4 using values of !P ˆ 26,000 cmÿ1 and !d ˆ 1140 cmÿ1 for a thin gold film and d ˆ 1 as the dielectric constant of vacuum. This figure also shows the wave vector for p-polarized light incident on the metal in vacuum or air (! ˆ ckph ) (dotted line). It is evident that there is not a point in the graph where the frequency and wave vector for surface plasmons and incident light attain the same value because

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10.1 Dis persion r el ations f or sur face pl as mo ns on a metal w ith !P ˆ 26,000 cmÿ1 at the air-solid interface (solid line); light propagating in air (dotted line); and light propagating through the prism coupler (long p P =p 2 are shown for comparison. The dashes). The horizontal lines for !P and ! P = 1 ‡ d which in air or vacuum is plasmon cut-off p frequency is equal to ! equal to !P = 2.

these lines do not intersect. The significance of this is that it is practically impossible to excite surface plasmons by directing a light beam on a metallic surface without the appropriate experimental setup that allows coupling of surface plasmons and the electric vector of incident light. Figure 10.1 shows that it is necessary to decrease the slope of ! ˆ ckph in order to find a line that can intersect the surface plasmon dispersion relation given by eqn 10.3. That is, the experimental setup up needed to excite surface plasmons needs to be such that the relationship between light frequency, !, and wave vector of the photons, kph, exists with a slope smaller than the speed of light in vacuum, c. This can be conveniently accomplished using the Kretschmann configuration (Kretschmann, 1971). This concept is illustrated in Fig. 10.2, where the wave vector of the photons within a prism of refractive index np becomes np
kph . In this figure, the wave vectors of the incident light within the prism and that of the surface plasmons are represented as arrow vectors. This arrangement comprises the operating component of all commercial SPR instruments. In the Kretschmann configuration, the light beam directed through the coupling prism with index of refraction, np, can be incident at a variable angle, . Thus, the wave vector of the photons incident on the metal film has a component parallel to the surface of the metal given by ! kx ˆ np sin  c

10:5

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10.2 Schematic representation of the experimental setup for excitation of surface plasmons using the Kretschmann configuration. At a specific angle of incidence the dispersion relation for surface plasmons and the horizontal component of the incident photons becomes identical, resulting in excitation of surface plasmons.

In the Kretschmann configuration, when the angle of incidence of the light,  is varied, a point is reached where the frequency and momentum of the horizontal component of the photons matches that of the surface plasmons. This is schematically illustrated in Fig. 10.2 as vector arrows for the photons, np
kph , and plasmons, ksp. Figure 10.1 also plots the wave vector of light inside the prism at an angle of 41.5ë (straight dashed line). There is a point where this line intersects the surface plasmon relation (solid line). From Fig. 10.1 and eqn 10.5 we infer that the surface plasmons can be excited by a variety of combinations of angle of incidence, , and frequency of the photons, !. The intersection in Fig. 10.1 was simulated to occur for a frequency of 15,800 cmÿ1 ( ˆ 632.8 nm), which is the commonly used He-Ne red laser. When excitation of plasmons by photons occurs, the photons transfer their energy to the free electrons of the metal. This manifests as a decreased intensity in the reflected light beam. Thus, when varying the angle of incidence, a plot of the reflected intensity shows a dip or minimum at the angle where the surface plasmons are excited. The reflected intensity increases for values of  above and below the excitation of the surface plasmons. The behavior of the reflectivity curve can be described by the Fresnel equations for a system where light is reflected from the different interfaces (e.g., prism-metal, metal-dielectric, and dielectric-dielectric interfaces). A web-based calculation program for reflectivity curves has been made conveniently available by Professor Robert M. Corn at the Department of Chemistry, University of Wisconsin-Madison (http://corndog.chem.wisc.edu/fresnel/fcform.html). This program allows calculation of reflectivity curves versus angle of incidence for four-phase systems (prism, thin metal film, first dielectric and second dielectric layers). A program allowing N-phase Fresnel calculations is also available (http:// corninfo.chem.wisc.edu/calculations.html). Using this program and literature values for the dielectric constants of gold and silver, the reflectivity curves of Fig. 10.3 were calculated assuming a BK-7 coupling prism (n ˆ 1:515), and a thin films (50 nm) of Au or Ag with a thin layer of a dielectric material with

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10.3 Reflectivity profiles for the excitation of surface plasmons in air with the Kretschmann configuration using thin films of gold or silver.

n ˆ 1:45 and thickness of 1 nm in air (n  1:0). It is observed that the position of the reflectance minimum depends on the optical properties of the materials. Similarly, if we calculate the reflectivity profiles for a single material (e.g., gold) but at different wavelengths, one would obtain different reflectance profiles because the dielectric constants of the metal would be different at different wavelengths. This implies that one should be able to implement SPR, in an alternative fashion, by using multiple wavelengths (e.g., from a wide spectrum light source or multiple lasers) and a given angle of incidence, to obtain reflectivity profiles where the reflectance at a particular wavelength goes through a minimum (Homola, 1998; Homola, 1999; Frutos, 1999). The use of SPR as a surface analytical technique lies in its surface sensitivity. If the dielectric constant in the vicinity of the thin metal film is perturbed, the reflectance profiles change. Figure 10.4 shows a simulation of reflectivity profiles for a thin gold film covered by dielectric material of varying thickness and n ˆ 1:45 in water (nH2 O ˆ 1:33). The position of the minimum changes with increased thickness of the dielectric layer and this change is sensitive to nanometer amounts of material. Thus, SPR is a powerful and very sensitive technique that allows detection of very small amounts of material depositing on the surface of a thin metal film. It also allows monitoring of changes in optical properties adjacent to a surface. In fact, SPR can perform very sensitive measurements of refractive index (down to 10ÿ6 refractive index units). Using SPR, it is possible to detect perturbations in optical surface properties by comparing reflectivity measurements before and after the change has occurred.

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10.4 Simulated reflectivity profiles for the excitation of surface plasmons in water (n ˆ 1:33) for a thin film of gold covered by a thin layer of a dielectric material with n ˆ 1:45 and varying thickness (0, 2, 5 and 8 nm). At the appropriate angle of incidence (dotted line), reflectance measurements can be performed to monitor fast changes in the vicinity of the metallic thin film induced by the adsorption of molecules. As the index of refraction increases in the vicinity of the thin metal film, the curves shift to the right and at the fixed angle indicated by the dotted line the reflectance increases.

Additionally, SPR also allows for very sensitive, real time, monitoring of these changes near a thin metal film. If we analyze Fig. 10.4, we can see that if we perform our reflectance measurements at a fixed angle in the steepest (and near linear) region of the reflectivity profile (dotted vertical line), the changes occurring in the vicinity of the surface (e.g., increased thickness of the dielectric layer) will shift the reflectivity profiles. These changes will increase the reflectivity at constant angle of incidence because the curves get progressively shifted from the minimum SPR angle. This is probably the most powerful feature of SPR that allowed it to become widely used for the determination of biospecific interactions. Due to the sensitive detection of changes in index of refraction, SPR is also very sensitive to fluctuations in temperature. Because of this, SPR measurements usually include a reference element to subtract contributions from thermal fluctuations or fluctuations on the bulk index of refraction. A reference element can be a section of the thin film where adsorption does not take place either because the surface does not allow adsorption of molecules (O'Brien, 1999; Lu, 2001; Akimoto, 2003; Boozer, 2003) or because it encompasses a microfluidics channel where the adsorbing molecules are not

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delivered (Karlsson, 1995). A second option is to include a temperature sensor to compensate for fluctuations in the index of refraction with temperature through appropriate numerical manipulation of the collected signal (MeleÂndez, 1996, 1997).

10.2.2 Metallic nanoparticles The optical properties of metallic colloids are also a result of the SPR phenomenon. Michael Faraday was the first to explain the colors of colloidal dispersions of gold and ruby glass in a qualitative theoretical framework that formed the foundations for more elaborate theories on the origin of colors of metallic nanoparticles (Schmitt, 1999; Kreibig, 1995; Bohren, 1983; Mulvaney, 1996). For metallic nanospheres that are much smaller than the wavelength of light (R  ) the optical properties of metallic clusters can be represented by a set of equations known as the Maxwell-Garnett formula (Schmitt, 1999). The dielectric constant of metals is represented by a complex number, which is related to the index of refraction as follows  ˆ 0 ‡ i00 ˆ …n ‡ ik†2

10:6

where n is the index of refraction and k is the absorption coefficient. Thus, the absorption coefficient can be expressed in terms of the real and imaginary components of the dielectric constant as r   1 p 02 ‡ 002 ÿ 0 kˆ 10:7 2 which is related to the extinction coefficient, ˆ 4k=. An expression for the overall absorption was derived first by Mie in 1908 by integrating over a medium filled with particles such that, 3=2

k ˆ 9=2d …N 4=3R3 †

00m

…0m ‡ 2d †2 ‡ 002 m

10:8

where N is the number of particles per unit volume. The term in parentheses represents the filling factor f ˆ N 4=3R3 or volume fraction of particles in the system. However, this expression is valid for dilute dispersions with noninteracting particles. If the filling factor is large, each particle is subjected to an average polarization field due to the surrounding particles. For an isolated small spherical particle of radius R in a static electric field, the polarizability, , is given by m ÿ d 10:9  ˆ 40 d R3 m ‡ 2d where 0 is the dielectric constant of vacuum, d the dielectric constant of the media the nanoparticles are embedded into, and m the dielectric constant of the

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metallic particle. Equation 10.9 implies that the polarizability of a metallic nanoparticle will be maximum when the real part of 1=…m ‡ 2d † is maximum. For nondilute systems with interacting particles, an effective dielectric constant can be calculated as eff

2N  1 ‡ 2f  30 d ˆ d ˆ d N 1ÿf 1ÿ 30 d 1‡

with

ˆ

 30 d V

10:10

where V ˆ 4=3R3 is the particle's volume. This equation is known as the Maxwell-Garnett formula. Because of the static approximation assumption, the effective optical properties become independent of particle size. An absorption coefficient, k, can then be calculated according to eqn 10.7, using the effective dielectric constants obtained according the Maxwell-Garnett formula (eqn 10.10). Figure 10.5 shows the simulated extinction coefficient of 10 nm gold nanoparticles embedded in media of varying dielectric constants (corresponding to varying indexes of refraction, n ˆ 1.33, 1.35, 1.37 and 1.40). Tabulated experimental values of the dielectric constants of gold were used to simulate the spectra (Lide, 1999). This figure illustrates that for colloidal metals, the SPR absorption peak is sensitive to perturbation in the local index of refraction. In

10.5 Simulated optical absorption spectra of gold nanoparticles embedded in media of varying dielectric constants with n ˆ 1.33, 1.35, 1.37 and 1.40. Similar to the Kretschmann configuration, monitoring absorbance changes at a fixed wavelength in the steepest part of the absorption spectra allows for binding kinetics to be determined.

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general, increases in n result in slight red shifts in the surface plasmon resonance (SPR) peak position and an increased extinction coefficient of the SPR absorption. Interactions of materials (e.g., protein adsorption) with the nanoparticle's surfaces can induce perturbations in the local dielectric constant and cause similar spectral changes. This can be used and has been exploited as an alternative SPR sensing mechanism (Eck et al., 2001; Nath, 2002). In fact, it is possible to monitor kinetics of adsorption by following changes in absorbance at a fixed wavelength in a procedure similar to reflectance measurements at a fixed angle of incidence. The optical properties of gold nanoparticles are also very sensitive to aggregation. For aggregated particles with a center-to-center separation smaller than their radius, the oscillations of the surface plasmons can couple among the aggregated particles. In the simplest case of two aggregated particles, two excitation modes can exist, parallel or perpendicular relative to the axis pair. Longitudinal in-phase oscillations will become significantly red shifted due to coupling of the surface plasmons, which lowers the vibrational energy with respect to that of isolated nanospheres. In aggregated nanoparticles additional resonances appear at longer wavelengths. The overall result of these changes in the absorption spectra is a red shift of the maximum absorbance peak and a broadening of spectral features towards the larger wavelengths. In fact, the optical absorption spectra can show broad bands that extend into the near infrared region depending on the size of the aggregates and the number of particles in the aggregate (Bohren, 1983; Quinten, 1986; Wiesner, 1989; Kreibig, 1995; Storhoff, 2000). This phenomenon has been exploited to obtain a semiquantitative estimate of the degree of nanoparticle aggregation (Weisbecker, 1996; Mayya, 1997; Aslan, 2002). Recently, DNA hybridization detection has been implemented through proper surface modification of gold nanoparticles (Elghanian, 1997; Storhoff, 1998 and 2000). Similar approaches have been exploited to determine a variety of phenomena that result either in perturbations of the refractive index or aggregation of the nanoparticles.

10.3 Surface functionalization 10.3.1 Surface modification of planar, metallic, thin films The extreme sensitivity of SPR to perturbations in the local dielectric constant of the media surrounding the metal film or nanoparticles makes it extremely useful as a surface analysis tool for thin film characterization and formation. SPR is routinely used to detect specific binding of molecules at thin metal films. However, SPR lacks molecular specificity because it measures only changes in optical properties (e.g., refractive index or thickness of adsorbed layers) and some regard it as a `mass detector' because it generates signals from material deposited on surfaces (although the detection is not based on mass but optical

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properties). For analytical chemistry applications it is necessary to confer SPR with specific molecular recognition capabilities. This is accomplished by tailoring the surface of metallic thin films such that only the target molecule elicits a change in the SPR signal. In biosensing and specific biomolecular interaction analysis applications this is accomplished through proper immobilization of biomolecules (usually antibodies or antigens but other approaches such as molecular imprinting of thin films can be used as well). The simplest method to attach antibodies or antigens to metallic surfaces is to exploit the natural tendency of proteins to adsorb by means of nonspecific interactions to most surfaces (Andrade, 1985; Horbett, 1987; Norde, 1991). This method is simple, rapid and effective but suffers from several limitations. Uncontrolled, nonspecific interactions between proteins and surfaces may lead to nonspecific protein adsorption (giving rise to false positives) and partial denaturation or loss of biological activity of the protein. The mere fact that nonspecific protein adsorption is a stochastic process will cause a fraction of the proteins to interact with the surface in an unfavorable orientation for binding to soluble analytes (e.g., when the antigen binding site in an antibody faces the surface, they will be unavailable for interaction with antigens even if the protein has not been denatured). Additionally, proteins adsorbed by means of nonspecific interactions are prone to be displaced by other proteins with larger surface affinity. The latter may result in a decreased response of the biosensor with time, increased nonspecific interactions, or both. Because of this, it is desirable to count with facile methods of protein immobilization that minimize the aforementioned problems. An ideal method would have to allow robust immobilization of the biomolecules. Such method must minimize nonspecific adsorption of proteins, be conducive to maintaining the biological activity of the proteins, allow uniform and optimal orientation of immobilized receptors, and should provide for robust attachment of the desired biomolecules such that displacement and/or degradation does not proceed to a significant extent. These requirements point to the need for well-defined surface chemistries that can be produced in a reproducible manner. The spontaneous formation of self-assembled monolayers (SAMs) on metals such as gold, silver and copper has been widely exploited to create surfaces with well-defined properties (Ulman, 1991). In particular, alkanethiol-based selfassembled monolayers are widely exploited to tailor surface properties because of the strong affinity of the thiol group for the aforementioned metals. By far the most widely used metal for thin film SPR implementation is gold, due to its relatively chemical inertness. In contrast to silver or copper, gold does not become covered with an oxide layer upon exposure to oxygen from air. This makes manipulation of these films more convenient for subsequent chemical functionalization procedures with alkanethiols. In contrast, silver and copper can readily form oxide layers that make difficult the formation of well-ordered SAMs unless precautionary measures are taken to avoid their exposure to

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oxygen. Thus, much higher quality SAMs are obtained using gold surfaces. A variety of surface functional groups (methyl, carboxyl, amine, hydroxyl, and others) can be introduced by forming SAMs on gold using x-functionalized alkanethiols. For the study of protein interactions with surfaces or bimolecular specific interaction analysis it is necessary to minimize nonspecific protein adsorption. Few surfaces can resist or prevent nonspecific interactions of proteins and the most widely used are surfaces presenting oligo(ethylene glycol) groups (PaleGrosdemange, 1991; Prime, 1993). Other surface chemistries have provided adequate resistance to nonspecific protein interactions (Luk, 2000) but their surface chemistry is more elaborate and molecular precursors to make these are not commercially available. Given that oligo(ethylene glycol) surface groups prevent nonspecific interactions of proteins with surfaces, they have been used as a platform to present ligands that interact with molecules in solution. With this approach, surfaces consisting of a specific molecule embedded within a nonfouling matrix are displayed on thin gold films to study specific interactions with ligand binding proteins using SPR. Some examples of extensively explored systems are peptide sequences (Roberts, 1998), benzenesulfonamide (Lahiri, 1999) and biotin (Haeussling, 1991; Spinke, 1993; Perez-Luna, 1999; Nelson, 2001). Biotin deserves a special attention because of its strong interaction with streptavidin and avidin. Streptavidin and avidin are tetrameric, biotin binding proteins that have four biotin binding pockets located in pairs at opposite sides of these proteins. Thus, they function as molecular adaptors for further functionalization. Avidin or streptavidin will attach to biotinylated surfaces almost irreversibly (Ka ~ 1013 M) so that they interact with the surface through two biotin groups, presenting the other pair of biotin binding pockets away from the surface. The latter can be used for further functionalization of surfaces since biotin can be used to functionalize proteins, nucleotides, and peptides. Other approaches for immobilization of proteins may involve the use of FLAGTM peptides (Wegner, 2002) or histidine groups (Sigal, 1996). Chemical moieties such as carboxyl, amine, anhydrides and photosensitive groups have also been explored to immobilize biomolecules. A recent review of these methods was presented by Yang in 1999. Lipid bilayers can also be implemented to effect immobilization of proteins onto thin gold films for SPR analysis (Plant, 1995). A further aspect of SPR that can be exploited in order to increase its sensitivity is the fact that a surface plasmon is a bound, nonradiative evanescent wave with maximum amplitude at the surface, and which decays exponentially into the dielectric (Hanken, 1996; Knoll, 1998). Thus, a surface plasmon (and thus the detection range of SPR) extends away from the surface to a distance on the order of a hundred nanometers. Surface immobilization of receptors (antibodies, ligands, nucleotides, peptides) on planar surfaces results in a layer of material with a thickness that is usually less than 10 nm. Thus, interactions

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with these surfaces could occur within an additional 5 nm above the metallic thin film. This implies that only a fraction of the sensing region of SPR is sampled with conventional 2-D immobilization of biomolecules. Further gains in sensitivity can be achieved if the immobilized molecules are distributed throughout the first 100±200 nm away from the metal surface. The sensing chips for the BIACore SPR instrument developed by Pharmacia (LoÈfaÊs, 1990; JoÈnsson, 1991) consist of carboxylated dextran chains attached to the surface such that they extend about 100 to 200 nm above the surface of the gold film. This layer allows for immobilization of antibodies and other biomolecules throughout the region sampled by the evanescent SPR field. Thus, it allows for a larger number of receptors (and therefore binding events) per unit area of metallic thin film than with 2-D planar monolayers. This is schematically depicted in Fig. 10.6. Further advantages of surface grafted dextran are facile introduction of chemistries amenable for immobilization of biomolecules (O'Shannessy, 1992), low nonspecific interactions because dextran is a hydrophilic biopolymer, and low mass diffusion resistance for analytes within this hydrogel because the volume fraction of water within this dextran layer can be more than 97% (LoÈfaÊs, 1990). Because dextran allows immobilization of more receptors, sensitivity is increased by one order of magnitude. In fact, low molecular analytes (MW < 1000 Da) can be detected using these layers, something that is not possible using conventional 2-D immobilizations procedures (Karlsson, 1995). The formation of alkanethiol based SAMs is by far the most commonly employed surface modification procedure for SPR based applications. In fact, a hydroxyl functionalized SAM is an intermediate step during the dextran grafting

10.6 Immobilization of antibodies using dextran layers allows for more receptors per unit area than is possible on a single monolayer.

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procedure for the sensing chips in the SPR BIACore instrument (JoÈnsson, 1991; LoÈfaÊs, 1990). However, additional approaches have been exploited such as the use of alkanethiol based SAMs (Haeussling, 1991; Spinke, 1993; PeÂrez-Luna, 1999; Nelson, 2001), lipid bilayers (Plant, 1995), and adsorption of polymers (Jordan, 1997) to name a few. Alkyl silanes are a robust and attractive option for surface functionalization. However, they do not form SAMs on gold surfaces; even so, it is possible to form thin films of SiO2 on gold by the sol-gel technique that allow further functionalization with silanes. If the SiO2 layer is small enough and well within the evanescent field of SPR, comparable sensitivity to SAMs of alkane thiols on gold is obtained with the additional advantage of a more robust and stable monolayer (Kambhampati, 2001).

10.3.2 Surface modification of metallic nanoparticles Surface modification of metallic nanoparticles can be implemented to take advantage of the SPR phenomena for applications similar to those involving thin metallic films. In principle, the same chemical manipulation procedures could be used. However, the surface modification of metallic colloid dispersions presents additional challenges. The stability of metallic nanoparticles in water is a result of electrostatic repulsion among charged particles. In the particular case of colloidal gold, the charged surfaces result from adsorption of anions. Surface modification procedures, such as those based on chemisorption of alkanethiols displace these charges with the resultant effect of decreased electrostatic repulsion. Often, this leads to flocculation and fusion of the metallic cores, an irreversible process that results in losing the nanoparticles to an aggregated state without the advantageous optical properties of individual nanoparticles. Some of the approaches used to circumvent this problem are the use of dilute solutions, dendrimers (Manna, 2001), polymers (Mangeney, 2002), chemisorption of alkanethiols at different ionic strengths and pH (Weisbecker, 1996; Mayya, 1997), or surface modification in the presence of stabilizing agents such as surfactants (Aslan, 2002). Additionally, if the metallic nanoparticles are immobilized on a solid support (monolayers of nanoparticles), further modification procedures can be employed without inducing their aggregation (Nath, 2002).

10.4 Applications 10.4.1 Surface plasmon resonance using planar, metallic, thin films The extremely high sensitivity of SPR to changes in surface optical properties (i.e., refractive index) makes it a useful tool in the characterization of thin films at metal surfaces. The high sensitivity of this technique has prompted its

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application to the study of self assembly of surface active molecules (Peterlinz, 1996), the process of formation of ultrathin multilayers (Cheng, 1999), adsorption of biomolecules (Mrksich, 1995), electrochemical processes such as surface oxidation (Hanken, 1996), the study of biomolecular interactions such as ligand-receptor, antigen-antibody, protein-protein, RNA-protein, cell-protein interactions, DNA hybridization, drug screening, and kinetics of gene expression. Its extremely high sensitivity to changes in refractive index has also prompted the application of SPR as an extremely sensitive temperaturemonitoring device (Chadwick, 1993). Surface plasmon resonance can be implemented as an imaging technique. In one variation of SPR imaging an expanded beam is directed over a large area of a thin metal film at a fixed angle (Hickel, 1990; Jordan, 1997; Thiel, 1997; Berger, 1998). If there are spatial variations of refractive index on the metal film surface these will appear with different reflected intensities. This approach has the advantage of allowing the imaging of a large area over a surface. It allows examination of microscopic features on a surface and in combination with 2-D patterning of biomolecules it can be used as a powerful tool for monitoring or screening of simultaneous binding events. However, this imaging mode can be sensitive to fluctuations in the illumination intensity. These effects can usually be subtracted from a non-interacting reference element. A second variation of SPR imaging consist on focusing a `wedge' or `fan' of light along a line on the surface of the metal film (O'Brien, 2001). In this modality, the reflected intensities are recorded using a charged coupled device (CCD) camera. This information then becomes a matrix of reflected intensities where the rows (or columns) contain the reflected intensities of a point on the surface at all the differing angles of incidence. Thus, the reflected wedge or fan of light will contain a dark band indicative of the plasmon resonances. This second imaging mode allows sampling of only a line along the surface. However, it permits accurate determination of the SPR angle and is less sensitive to fluctuations in the illumination source because the image is acquired at once. Thus, any fluctuations in illumination intensity are recorded simultaneously and the position of the angle on minimum intensity remains unchanged. The latter may be advantageous for kinetic determinations of simultaneous binding events on a surface.

10.4.2 Surface plasmon resonance using metallic nanoparticles Metallic colloids also exhibit a strong SPR absorption band in their spectra. However, it has not been until recently that these effects are being explored for the study of biomolecular specific interactions and surface phenomena at solidliquid interfaces. For metallic nanoparticles, the characteristics of the SPR band are affected not only by changes in refractive index in the vicinity of the metallic

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10.7 Schematic representation of the concept of nanoparticle aggregation for colorimetric detection (based on SPR) of DNA, transition temperatures of elastin peptides and divalent heavy metals (from J. Am. Chem. Soc. 122(19), (2000), 4640±4650, J. Am. Chem. Soc., 123(34), (2001), 8197±8202, and J. Phys. Chem. B.106(18), (2002), 4647±4651).

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surface, but also by the degree of interaction of the nanoparticles. The latter possess both, an additional advantage but also a challenge for the implementation of metallic nanoparticles in the detection of surface phenomena. As mentioned earlier, the degree of interaction (aggregation) of these nanoparticles may give rise to additional resonances appearing at lower energies (longer wavelengths) upon aggregation. This has been implemented to great advantage for colorimetric detection of DNA hybridization (Elghanian, 1997; Storhoff, 1998 and 2002), protein-induced aggregation of ligand-functionalized colloids (Otsuka, 2001) and antigen-functionalized colloids (Thanh, 2002), for the study of thermodynamic transitions of thermoresponsive polymers (Nath, 2001), as a detection method for divalent cations such as Pb+2 (Kim, 2001; Berchmans, 2002), to evaluate the stability of colloidal metals during their manipulation (Weisbecker, 1996; Mayya, 1997; Aslan, 2002). Examples of these transduction schemes are shown in Fig. 10.7. Even in the absence of aggregation metallic nanoparticles are sensitive to local perturbations in the refractive index and this has also been exploited to detect adsorption of macromolecules. A scheme that offers great potential to perform SPR sensing with a minimum investment consists on metallic nanoparticles immobilized on a solid support. These will also be sensitive to local changes in refractive index, but they would not aggregate because they are affixed to a solid surface. However, even if they do not have the capability to aggregate, they can be used to detect binding of molecules on their surface. This has recently been implemented for the study of biomolecular specific interactions and holds enormous potential for further facile implementation of SPR sensing and imaging using nanoparticles (Nath, 2002).

10.5 Conclusion The use of SPR as a detection technique is reaching a mature stage for conventional biomolecular interaction analysis. Surface immobilization of receptors and other biological molecules on sensing chips is now routinely performed with automated microfluidics systems in commercial instruments and surface modified sensing chips are commercially available with dextran layers, biotin groups, and other ready-to-use surface functionalities. Due to its extreme sensitivity and capability to perform reagentless detection in real time, SPR will continue to play a major role as a surface analytical technique for a large number of applications involving the formation of ultra thin films and the adsorption of macromolecules on metallic surfaces. Imaging systems based on the SPR technique also have great potential as high-throughput screening tools in medicine and combinatorial chemistry. However, this has not been exploited to the same extent as single spot SPR sensing. Although commercial SPR imaging systems exist, delivering a totally automated SPR imaging system requiring only minimum intervention from the end user is more difficult than single spot

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sensing. Additionally, 2-D patterning of receptors (Yang, 1999) and obtaining the molecules to be patterned require additional expertise that may deter widespread use of this technique. The enormous potential of sensitivity and the large amounts of information this technique could provide remains largely underexploited. In conventional SPR of materials with very low or zero absorption coefficients, detection is based mostly on changes in the real part of the refractive index. Recently, encouraging attempts at using SPR with molecules with significant light absorption have prompted additional research that may further enhance the sensitivity of detection (Kurihara, 2002a, 2002b). In a limited, but significant number of cases, this new mode of SPR detection offers additional advantages when conventional SPR gives only marginal responses (e.g., with molecules of small molecular weight). Kurihara et al. (2002b) have shown, in a numerical and experimental model system, how materials with large absorption coefficients can greatly enhance the detection limits of this technique. In fact, by using a thin sodium ion-selective membrane containing a sensitive ionophore, it has been demonstrated that this technique can achieve detection of sodium ions in the 10ÿ6 to 10ÿ1 M range. It is expected that novel transduction schemes coupled to the high sensitivity of absorption based SPR will make possible real time, reagentless detection of small molecules in the near future.

10.6 Acknowledgements Financial support provided by the Armour College of Engineering, Department of Chemical and Environmental Engineering, and the Graduate College at the Illinois Institute of Technology are gratefully acknowledged.

10.7 References Akimoto T., Ikebukuro K. and Karube I., (2003) `A Surface Plasmon Resonance Probe with a Novel Integrated Reference Sensor Surface', Biosens. Bioelectr., 18, 1447±1453. Andrade J. D., (1985), `Principles of Protein Adsorption', in: Surface and Interfacial Aspects of Biomedical Polymers, J. Andrade, Ed. Plenum Publ., New York, 1±80. Aslan K. and PeÂrez-Luna V.H., (2002) `Surface Modification of Colloidal Gold by Chemisorption of Alkanethiols in the Presence of a Nonionic Surfactant', Langmuir, 18, 6059±6065. Berchmans S., Thomas P. J. and Rao C. N. R., (2002) `Novel Effects of Metal Ion Chelation on the Properties of Lipoic Acid-Capped Ag and Au Nanoparticles', J. Phys. Chem. B., 106, 4647±4651. Berger C. E. H., Beumer T. A. M., Kooyman R. P. H. and Greve J., (1998) `Surface Plasmon Resonance Multisensing', Anal. Chem., 70, 703±706. Bohren C. F. and Huffman D.R., (1983) Absorption and Scattering of Light by Small Particles, Wiley, New York.

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Boozer C., Yu Q. M., Chen S. F., Lee C. Y., Homola J., Yee S. S. and Jiang S.Y., (2003) `Surface Functionalization for Self-Referencing Surface Plasmon Resonance (SPR) Biosensors by Multi-Step Self-Assembly', Sens. Actuators B-Chem., 90, 22±30. Chadwick B. and Gal M., (1993) `An Optical Temperature Sensor Using Surface Plasmons', Jpn. J. Appl. Phys., 32, 2716±2717. Cheng Y. and Corn R. M., (1999) `Ultrathin Polypeptide Multilayer Films for the Fabrication of Model Liquid/Liquid Electrochemical Interfaces', J. Phys. Chem. B., 103, 8726±8731. Eck D., Helm C. A., Wagner N. J. and Vaynberg K. A., (2001) `Plasmon Resonance Measurements of the Adsorption and Adsorption Kinetics of a Biopolymer onto Gold Nanocolloids', Langmuir, 17, 957±960. Elghanian R., Storhoff J. J., Mucic R. C., Letsinger R. L. and Mirkin C. A., (1997) `Selective Colorimetric Detection of Polynucleotides Based on the DistanceDependent Optical Properties of Gold Nanoparticles', Science 277, 1078±1081. Fano U., (1941) J. Opt. Soc. Am., 31, 213±222. Faraday M., (1857) `The Bakerian Lecture: Experimental Relations of Gold (and Other Metals) to Light, Philos. Trans. R. Soc. London, 147, 145±181. Feldman L.C. and Mayer J. W. (1986) Fundamentals of Surface and Thin Film Analysis, North Holland, New York. Flanagan M.T. and Pantell R.H., (1984) `Surface Plasmon Resonance and Immunosensors', Electronics Lett., 20, 968±970. Frutos A. G., Weibel S. C. and Corn R. M., (1999) `Near-Infrared Surface Plasmon Resonance Measurements of Ultrathin Films. 2. Fourier Transform SPR Spectroscopy', Anal. Chem., 71, 3935±3940. Haeussling L., Ringsdorf H., Schmitt F. J. and Knoll W., (1991) `Biotin-Functionalized Self-Assembled Monolayers on Gold: Surface Plasmon Optical Studies of Specific Recognition Reactions' Langmuir, 7, 1837±1840. Hanken D.G., Jordan C. E., Frey B. L. and Corn R. M., (1996) `Surface Plasmon Resonance Measurements of Ultrathin Organic Films at Electrode Surfaces', In: Electroanalytical Chemistry a Series of Advances, Volume 20, Marcel Dekker, Inc., New York. Hickel W. and Knoll W. J., (1990) `Surface Plasmon Optical Characterization of Lipid Monolayers at 5 mm Lateral Resolution', Appl. Phys. 67, 3572±3575. Homola J. and Yee S. S., (1998) `Novel Polarization Control Scheme for Spectral Surface Plasmon Resonance Sensors', Sensors and Actuators B, 51, 331±339. Homola J., Koudela I. and Yee S. S., (1999) `Surface Plasmon Resonance Sensors Based on Diffraction Gratings and Prism Couplers: Sensitivity Comparison', Sensors and Actuators B, 54, 16±24. Horbett T. A. and Brash J. L., (1987) `Proteins at Interfaces: Current Issues and Future Prospects', in: Proteins at Interfaces: Physicochemical and Biochemical Studies, T. A. Horbett and J. L. Brash, eds, ACS Symposium Series, 343, American Chemical Society, Washington, DC, 343, 1±33. JoÈnsson U., FaÈgerstam L., Ivarsson B., Johnsson B., Karlsson R., Lundh K., LoÈfaÊs S., Persson B., Roos H., RoÈnnberg I., SjoÈlander S., Stenberg E., StaÊhlberg R., È stlin H. and Malmqvist M., (1991) `Real-Time Biospecific Interaction Urbaniczky, O Analysis Using Surface Plasmon Resonance and a Sensor Chip Technology', BioTechniques, 11, 620±627. Jordan C. E. and Corn R. M., (1997) `Surface Plasmon Resonance Imaging

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Measurements of Electrostatic Biopolymer Adsorption onto Chemically Modified Gold Surfaces', Anal. Chem., 69, 1449±1456. Kambhampati D. K., Jakob T. A. M., Robertson J. W., Cai M., Pemberton J. E. and Knoll W., (2001) `Novel Silicon Dioxide Sol-Gel Films for Potential Sensor Applications: a Surface Plasmon Resonance Study' Langmuir, 17, 1169±1175. Karlsson R. and StaÊhlberg R., (1995) `Surface Plasmon Resonance Detection and Multispot Sensing for Direct Monitoring of Interactions Involving Low-MolecularWeight Analytes and for Determination of Low Affinities', Anal. Biochem., 228, 274±280. Kim Y., Johnson R. C. and Hupp J. T., (2001) `Gold Nanoparticle-Based Sensing of ``Spectroscopically Silent'' Heavy Metal Ions', Nano Lett., 1, 165±167. Kittel C., (1956) Introduction to Solid State Physics, 2nd edn, Wiley, New York. Knoll W., (1998) `Interfaces and Thin Films as Seen by Bound Electromagnetic Waves', Annu. Rev. Phys. Chem., 49, 569±638. Kooyman R. P. H., Kolkman H., Van Gent J. and Greve J., (1988) `Surface Plasmon Resonance Immunosensors: Sensitivity Considerations', Anal. Chim. Acta, 213, 35± 45. Kreibig U. and Vollmer M., (1995) Optical Properties of Metal Clusters, SpringerVerlag, Berlin. Kretschmann E., (1971) Z. Phys., B241, 313±324. Kurihara K. and Suzuki K., (2002a) `Theoretical Understanding of an Absorption-Based Surface Plasmon Resonance Sensor Based on Kretchmann's Theory', Anal. Chem., 74, 696±701. Kurihara K., Nakamura K., Hirayama E. and Suzuki K., (2002b) `An Absorption-Based Surface Plasmon Resonance Sensor Applied to Sodium Ion Sensing Based on an Ion-Selective Optode Membrane', Anal. Chem.,74, 6323±6333. Lahiri J., Isaacs L., Grzybowski B., Carbeck J. D. and Whitesides, G. M., (1999) `Biospecific Binding of Carbonic Anhydrase to Mixed SAMs Presenting Benzenesulfonamide Ligands: A Model System for Studying Lateral Steric Effects', Langmuir, 15, 7186±7198. Leidberg B., Nylander C. and LunstroÈm I., (1983) `Surface Plasmon Resonance for Gas Detection and Biosensing', Sensors and Actuators, 4, 299±304. Lide, D. R., ed., (1999) CRC Handbook of Chemistry and Physics, 80th edn, CRC Press, Boca Raton. LoÈfaÊs S. and Johnsson B., (1990) `A Novel Hydrogel Matrix on Gold Surfaces in Surface Plasmon Resonance Sensors for Fast and Efficient Covalent Immobilization of Ligands', J. Chem. Soc., Chem. Commun. 1526±1528. Lu H. B., Homola J., Campbell C. T., Nenninger G. G., Yee S. S. and Ratner B. D., (2001) `Protein Contact Printing for a Surface Plasmon Resonance Biosensor with On-Chip Referencing' Sens. Actuators B-Chem., 74, 91±99. Luk Y.-Y., Kato M. and Mrksich M., (2000) `Self-Assembled Monolayers of Alkanethiolates Presenting Mannitol Groups are Inert to Protein Adsorption and Cell Attachment' Langmuir, 16, 9604±9608. Mangeney C., Ferrage F., Aujard I., Marchi-Artzner V., Jullien L., Ouari O., Rekai E. D., Laschewsky A., Vikholm I. and Sadowski J. W., (2002) `Synthesis and Properties of Water-Soluble Gold Colloids Covalently Derivatized with Neutral Polymer Monolayers', J. Am. Chem. Soc., 124, 5811±5821. Manna A., Imae T., Aoi K., Okada M. and Yogo T., (2001) `Synthesis of Dendrimer-

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Passivated Noble Metal Nanoparticles in a Polar Medium: Comparison of Size between Silver and Gold Particles', Chem. Mater., 13, 1674±1681. Mayya K. S., Patil V. and Sastry M., (1997) `On the Stability of Carboxylic Acid Derivatized Gold Colloidal Particles: The Role of Colloidal Solution pH Studied by Optical Absorption Spectroscopy', Langmuir, 13, 3944±3947. MeleÂndez J., Carr R., Bartholomew D.U., Kukanskis K., Elkind J., Yee S., Furlong C. and Woodbury R., (1996) `A Commercial Solution for Surface Plasmon Resonance Sensing', Sensors and Actuators B, 35, 1±5. MeleÂndez J., Carr R., Bartholomew D., Taneja H., Yee S., Jung C. and Furlong C., (1997) `Development of a Surface Plasmon Resonance Sensor for Commercial Applications', Sensors and Actuators B, 38±39, 375±379. Mie G., (1908) `BeitraÈge zur Optik truÈber Medien speziell kolloidaler MetalloÈsungen' Ann. Phys. 25, 377±445. Mrksich M., Sigal G. B. and Whitesides G. M., (1995) `Surface Plasmon Resonance Permits in situ Measurement of Protein Adsorption on Self-Assembled Monolayers of Alkanethiolates on Gold', Langmuir, 11, 4383±4385. Mulvaney P., (1996) `Surface Plasmon Spectroscopy of Nanosized Metal Particles', Langmuir, 12, 788±800. Nath N. and Chilkoti A., (2001) `Interfacial Phase Transition of an Environmentally Responsive Elastin Biopolymer Adsorbed on Functionalized Gold Nanoparticles Studied by Colloidal Surface Plasmon Resonance', J. Am. Chem. Soc., 123, 8197± 8202. Nath N. and Chilkoti A., (2002) `A Colorimetric Gold Nanoparticle Sensor to Interrogate Biomolecular Interactions in Real Time on a Surface', Anal. Chem. 74, 504±509. Nelson K. E., Gamble L., Jung L. S., Boeckl M. S., Naeemi E., Golledge S. L., Sasaki T., Castner D. G., Campbell C. T. and Stayton P. S., (2001) `Surface Characterization of Mixed Self-Assembled Monolayers Designed for Streptavidin Immobilization' Langmuir, 17, 2807±2816. Norde W. and Lyklema J., (1991) `Why Proteins Prefer Interfaces', J. Biomater. Sci.: Polymer edn, 2, 183±202. O'Brien M. J., Brueck S.R.J., PeÂrez-Luna V. H., Tender L. M. and LoÂpez G. P., (1999) `SPR Biosensors: Simultaneously Removing Thermal and Bulk-Composition Effects', Biosens. Bioelectr., 14, 145±154. O'Brien M. J., Brueck S. R. J., PeÂrez-Luna V. H. and LoÂpez G. P., (2001) `A Surface Plasmon Resonance Array Biosensor Based on Spectroscopic Imaging', Biosensors and Bioelectronics, 16, 97±108. O'Shannessy D. J., Brigham-Burke M. and Peck K., (1992) `Immobilization Chemistries Suitable for Use in the BIAcore Surface Plasmon Resonance Detector', Anal. Biochem., 205, 132±136. Otsuka H., Akiyama Y., Nagasaki Y. and Kataoka, K., (2001) `Quantitative and Reversible Lectin-Induced Association of Gold Nanoparticles Modified with Lactosyl±mercapto-poly(ethylene glycol)', J. Am. Chem. Soc., 123, 8226±8230. Pale-Grosdemange C., Simon E. S., Prime K. L. and Whitesides G. M., (1991) `Formation of Self-Assembled Monolayers by Chemisorption of Derivatives of Oligo(ethylene glycol) of Structure HS(CH2)11(OCH2CH2)mOH on gold', J. Am. Chem. Soc., 113, 12±20. PeÂrez-Luna V. H., O'Brien M. J., Opperman K. A., Hampton P. D., Lopez G. P., Klumb L. A. and Stayton P. S., (1999) `Molecular Recognition between Genetically

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Engineered Streptavidin and Surface-Bound Biotin' J. Am. Chem. Soc., 121, 6469± 6478. PeÂrez-Luna V. H., Aslan K. and Betala P., (2004) `Colloidal Gold', in: Encyclopedia of Nanoscience and Nanotechnology. H. S. Nalwa, ed., American Scientific Publishers. Peterlinz K. A. and Georgiadis R., (1996) `In Situ Kinetics of Self-Assembly by Surface Plasmon Resonance Spectroscopy', Langmuir, 12, 4731±4740. Plant A.L., Brigham-Burke M., Petrella E.C. and O'Shannessy D. J., (1995) `Phospholipid/Alkanethiol Bilayers for Cell Surface Receptor Studies by Surface Plasmon Resonance', Anal. Biochem. 226, 342±348. Powell C.J. and Swan J.B., (1959) `Origin of the Characteristic Electron Energy Losses in Aluminum', Phys. Rev., 115, 869±875. Prime K. L. and Whitesides G. M., (1993) `Adsorption of Proteins onto Surfaces Containing End-Attached Oligo(ethylene Oxide): a Model System using SelfAssembled Monolayers' J. Am. Chem. Soc., 115, 10714±10721. Quinten M. and Kreibig U., (1986) `Optical Properties of Aggregates of Small Metal Particles', Surface Sci. 172, 557±577. Rich R. L. and Myszka, D. G., (2001) `Survey of the Year 2000 Commercial Optical Biosensor Literature', J. Mol. Recognit.,14, 273±294. Ritchie R.H., (1957) `Plasma Losses by Fast Electrons in Thin Films', Phys. Rev., 106, 874±881. Roberts C., Chen C. S., Mrksich M., Martichonok V., Ingber D. E. and Whitesides, G. M., (1998) `Using Mixed Self-Assembled Monolayers Presenting RGD and (EG)3OH Groups to Characterize Long-Term Attachment of Bovine Capillary Endothelial Cells to Surfaces', J. Am. Chem. Soc., 120, 6548±6555. Schmitt J., MaÈchtle P. M., Eck D., MoÈwald H. and Helm C. A., (1999) `Preparation and Optical Properties of Colloidal Gold Monolayers', Langmuir, 15, 3256±3266. Sigal G. B., Bamdad C., Barberis A., Strominger J. and Whitesides G. M., (1996) `A SelfAssembled Monolayer for the Binding and Study of Histidine-Tagged Proteins by Surface Plasmon Resonance' Anal. Chem., 68, 490±497. Spinke J., Liley M., Schmitt F. J., Guder H. J., Angermaier L. and Knoll W., (1993) `Molecular Recognition at a Self-Assembled Monolayers: Optimization of Surface Functionalization', J. Chem. Phys. 99, 7012±7019. Stenberg E., Persson B., Roos H. and Urbaniczky C., (1991) `Quantitative Determination of Surface Concentration of Protein with Surface Plasmon Resonance Using Radiolabeled Proteins', J. Coll. Inerf. Sci., 141, 511±526. Stern E.A. and Farrell R.A., (1960) `Surface Plasma Oscillations of a Degenerate Electron Gas', Phys. Rev., 120, 130±136. Storhoff J. J., Elghanian R., Mucic R. C., Mirkin C. A. and Letsinger R. L., (1998) `OnePot Colorimetric Differentiation of Polynucleotides with Single Base Imperfections Using Gold Nanoparticle Probes' J. Am. Chem. Soc., 120, 1959±1964. Storhoff J. J., Lazarides A. A., Mucic R. C., Mirkin C. A., Letsinger R. L. and Schatz G. C., (2000) `What Controls the Optical Properties of DNA-Linked Gold Nanoparticle Assemblies?' J. Am. Chem. Soc., 122, 4640±4650. Thanh N. T. K. and Rosenzweig, Z., (2002) `Development of an Aggregation-Based Immunoassay for Anti-Protein A Using Gold Nanoparticles', Anal. Chem., 74, 1624±1628. Thiel, A. J.; Frutos, A. G.; Jordan, C. E.; Corn, R. M.; Smith, L. M., (1997) `In Situ Surface Plasmon Resonance Imaging Detection of DNA Hybridization to

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11

Ellipsometry for optical surface study applications Y M G E B R E M I C H A E L AND K T V G R A T T A N , City University, London

11.1 Introduction This chapter provides a brief discussion of the principles and practice of ellipsometric measurements for optical surface studies, especially using optical fibre-based devices. With the developments over recent years in the optical fibre communications market, a range of components and devices have been created and used in this field with the result that the `spin off' into optical fibre measurements and sensor systems has been very significant. This offers the possibility of the creation of a range of new devices employing technologies often adopted from the telecommunications field. Specialised optical fibres, in particular used outside the telecommunications industry, are now readily available and as a result highly birefringent (HiBi) fibre can be used to enable polarisation states to be maintained within fibres ± a major advantage for the development of compact, fibre optic ellipsometric instrumentation. This chapter reviews several instrumental developments based on fibre optic technology, and following a description of the underlying mathematical basis, the principles and applications of a range of ellipsometric devices are considered, focusing in particular upon polarisation based ellipsometry and fibre optic devices in particular. Reference is made to a range of techniques and devices reported in the literature and applications of sensor and measurement systems of this type are considered.

11.1.1 Polarisation of light and ellipsometry Light waves, being transverse electromagnetic waves are characterised by timevarying electric and magnetic field vectors, which are perpendicular to each other and normal to the wave propagation direction. The orientational characteristics of one of these carriers of the field, observed at a fixed point in space and time, defines the polarisation of the electromagnetic wave.1,2 For optical study purposes, it is conventional to retain the electric field strength, E, based on the fact that when light interacts with matter, the force exerted on the

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electrons by the electric field is much greater than that exerted by the magnetic field of the wave.3 Most of the light encountered every day is a chaotic mixture of light waves vibrating in all directions ± this is known as unpolarised light.4 Some light sources generate a polarised beam but in general, however, polarised light is obtained by passing the light beam through a polarising optical element or by causing the beam to make a reflection at a specific angle, called the Brewster angle.5 A plane polarised light beam (with a constant plane of vibration) can be resolved into two orthogonal optical disturbances, Ex and Ey, which are resultant components along the x and y axes.5 This may be expressed as: Ex …z; t† ˆ Ex0 ej…!tÿ2Z=†

11:1

Ey …z; t† ˆ Ey0 ej…!tÿ2Z=‡†

11:2

where  is the relative phase difference between the two waves, Ex and Ey, both of which are travelling in the z direction. Ex0 and Ey0 are the amplitudes of the x and y vibrations at z ˆ 0, ! is the angular frequency and  is the wavelength. If the two waves are considered to have the same phase, i.e.,  is zero or an integral multiple of 2, the resultant wave is a linearly polarised light beam with a fixed amplitude and a constant polarisation angle. Another case of particular interest is when both the component waves are 90ë out of phase and have equal amplitudes. When these two waves are combined, the tip of the resultant electric vector traces a circle, and the light beam is referred to as being circularly polarised. If the phase difference is other than 0ë or 90ë, the resultant electric vector will both rotate and also change its magnitude as well. The end point of the electric vector will trace out an ellipse, defining it as elliptically polarised light. In practice, there are several ways to obtain elliptically polarised light. Of primary interest in polarimetry is the fact that when linearly polarised light interacts with an optical surface, there is a change in amplitude as well as a shift in the phases of both the normal polarisation states. This change, in general, is not the same for both components; hence the resultant light will be elliptically polarised. Such polarised light carries valuable information about the interaction of light with matter and the various physical parameters which have been acting upon it.6 A variety of physical phenomena influence the state of polarisation of light. These include chemical interactions, molecular structures and mechanical stress which all impose changes in the polarisation state of an optical beam, and thus applications relying on the study of the polarisation state of light cover a vast range of engineering, medical, physical, agricultural and industrial processes.

11.1.2 Ellipsometry When linearly polarised light, of a known orientation, is reflected at an oblique incidence from a surface, then the reflected light is elliptically polarised. The

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shape and orientation of the ellipse and thus the result of the interaction of the light beam with the surface depends on the angle of incidence, the direction of the polarisation of the incident light, and the reflective properties of the surface. Ellipsometric measurement3,6±8 is a non-invasive and sensitive technique for determining the optical properties of surfaces and thin films by measuring the changes in the polarisation state of light when it is reflected from or transmitted through a sample. If the sample undergoes a change, as for example occurs when a thin film on a surface changes its thickness, then its reflection properties will also change. The ability to monitor these changes by measuring the state of the polarisation of the light, before and after the interaction with the substrate, permits an investigation of the processes occurring at the substrate or a characterising of the optical properties of the sample, illustrated schematically in Fig. 11.1. The most important application of ellipsometry in physical measurements is to study thin films and surfaces, and as a result this provides a means to study (and design) devices which contain them. In the context of ellipsometry, a thin film is one that ranges from essentially zero to several hundred nanometers in thickness. The sensitivity of an ellipsometer is such that a change of film thickness of a few tenths of a nanometer is usually relatively easy to detect by optical means. The process of ellipsometry requires a measurement of the change of the state of polarisation of light. This change involves the amplitude of the two orthogonal polarisation components and/or a change in their relative phase difference. These parameters, which are fundamental to the ellipsometric measurement cannot, however, be measured instantaneously. The reason for this is that the locus of the state of polarisation (SOP) goes through one period in a very short time, typically of the order of 1012 cycles per second.9 This instantaneous change and the random statistical nature of the change of the SOP of the unpolarised light is too fast for most practical measurements. Timeaveraged observations overcome the above measurement problem in practical applications and are satisfactory for most situations. Such an averaged measurement representation of polarised, partially polarised or unpolarised

11.1 The geometry of reflection ellipsometry. The change in polarisation upon reflection carries information on the characteristics of the sample.

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light involves using the Stokes parameters.9 These parameters are measurable quantities and are useful in analysing the evolution of the state of polarisation of the light on passing through polarising optical components. Thus, a partially polarised light beam can be represented mathematically by a 4  1 column vector whose elements are termed the four Stokes parameters. Thus if a beam of light passes through a succession of optical devices, the measured Stokes parameters at the output characterise the effect of the devices on the SOP of the input beam, i.e., Sout ˆ MSin , where Sin and Sout are the Stokes vectors of the input and output beams respectively and, M is a 4  4 matrix, which is characteristic of the optical device and the orientation of its optical axis. Such a representation is called the Muller matrix.1,5,9 The use of Jones calculus5,9 represents an approach to dealing with understanding the propagation of polarised light through polarising optical components. This involves a simpler 2  2-matrix formulation for optical polarising component representation. Jones calculus is applicable to polarised light with 100% degree of polarisation (DOP) only, otherwise the Muller matrix formulation should be selected.

11.1.3 Interaction of light with matter Reflection by ambient-substrate boundary system When a beam of light is incident at a boundary between two isotropic homogeneous media, with respective refractive indices, N0 and N1 as shown in Fig. 11.2, some of the light is reflected at the boundary and some of it is transmitted through the sample. The light can be resolved into two orthogonal components with the direction of vibration parallel and perpendicular to the plane of incidence, where the plane of incidence is defined as the plane normal to the sample containing both the incident and reflected beams. The ratio of the reflected electric vector to the incident wave is defined by the Fresnel reflection or transmission coefficients.5,10,11 These are given by: Erp N1 cos 0 ÿ N0 cos 1 ˆ rp ˆ Eip N1 cos 0 ‡ N0 cos 1

11:3

Ers N0 cos 0 ÿ N1 cos 1 ˆ rs ˆ Eis N0 cos 0 ‡ N1 cos 1

11:4

where the subscripts p and s refer to the waves parallel and perpendicular to the plane of incidence and Eip and Eis refer to the complex amplitudes of the incident electric vector while Erp and Ers are the corresponding reflected components. N0 and N1 are the refractive indices of the ambient and substrate system respectively and 1 and 2 are the angles of incidence and refraction respectively. The Fresnel equations provide a basis for examining the effect of reflection and refraction of polarised light at a boundary. It is convenient to express eqns

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11.2 Reflection and refraction of light at a boundary between the media.

11.3 and 11.4 in exponential form containing an amplitude and phase representation: rp ˆ jrp jejrp

11:5

rs ˆ jrs jejrs

11:6

where jrp j and jrs j are the respective ratios of the amplitudes of the reflected sinusoidal waves to those of the incident wave with s and p polarisations. p and s give the respective phase changes after reflection. The changes in either the amplitude or phase (or both) of the p and s polarisation states after reflection lead to the determination of the ellipsometric parameters and
as a ratio of the complex Fresnel reflection coefficients, rp and rs, such that: rp 11:7  ˆ  tan
ei
rs jrp j jrs j

11:8


ˆ rp ÿ rs

11:9

tan ˆ

where
is the phase change between the p and s polarisations and tan ( ) is the amplitude ratio. The value of
varies between 0ë and 360ë while ranges from 0ë to 90ë. Thus and
quantify the differential changes in amplitude and phase respectively of the incident s and p polarisations upon reflection from an optical sample. As the measurement technique involves the determination of both phase changes and amplitude ratios, it is a very sensitive technique offering accurate and repeatable ellipsometric values. The two measured parameters, and
are related to the polarisation change of the light caused through its interaction with a sample. A knowledge of and
allows the determination of the optical properties of a sample, such as its refractive index, thickness, the composition of a substrate and over layers, through the use of various equations and numerical algorithms, to produce a model describing the interaction of the light with matter. In an ambient-substrate system, the measurement of the complex refractive index, N1, of the substrate is usually of considerable interest. With a single

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ellipsometric measurement of and
, the complex refractive index can be readily computed. If a substitution is made for rp and rs in eqn 11.7, with their values from eqns 11.3 and 11.4 and using Snell's law,3 an equation for N1 can be obtained3 such that: " #1=2 4 2 sin …0 † 11:10 N1 ˆ N0 tan…0 † 1 ÿ …1 ‡ †2 where N0 is the ambient refractive index, 0 is the angle of incidence, and  is the measured complex reflectance ratio, given by tan… †ej
. In practice, since the measured quantity is the intensity of the reflected light, the reflectance measurements correspond to the square of the Fresnel reflection coefficients rp and rs. Thus, s r jrp2 j Rp ˆ : 11:11 ˆ jrs2 j Rs Reflection by ambient-film-substrate system The most common application of ellipsometry is in the analysis of thin films. This application is particularly useful, as the measurement technique is nondestructive and thus very thin films can be measured with no structural damage to the substrate and with high accuracy. Unlike the ambient-substrate system, in which only the phase and amplitude changes of the reflected polarisation are measured, in the ambient-film-substrate system an additional parameter is considered. This corresponds to the phase difference of the light rays traversing the film under study.4 As shown in Fig. 11.3, when the light beam is incident at the boundary between the ambient and the film media, some of the light is reflected back into the ambient medium and some is transmitted through the film and is, in turn, internally reflected back to the film at the film-substrate interface, with some

11.3 Reflection and transmission of a plane wave by an ambient-film-substrate system. Here d is the film thickness; 0 , the angle of incidence at the ambientfilm boundary and 1 and 2 , angles of refraction at the film and substrate interface.

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being refracted back to the ambient. If the Fresnel reflection coefficients at the ambient-film and film-substrate interfaces are denoted by r01 and r12, respectively for the s and p polarisation states and is the phase change the multi-reflected wave inside the film experiences as it traverses the film, the overall complex amplitude reflection coefficients, Rp and Rs, for the p and s polarisation states can be formulated by adding the successive partial waves that make up the resultant reflected wave.3 Thus Rp and Rs may be given by: Rp ˆ

r01p ‡ r12p eÿj2 1 ‡ r01p r12p eÿj2

11:12

Rs ˆ

r01s ‡ r12s eÿj2 1 ‡ r01s r12s eÿj2

11:13

where the film phase thickness, , is given by   d ˆ 2 …N12 ÿ N02 sin2 …0 †1=2 

11:14

and 0 is the angle of incidence, d is film thickness, N1 (ˆ n ÿ jk) is the complex refractive index of the film and  is the wavelength of the light. If  is expressed in terms of Rp and Rs (in eqns 11.12 and 11.13 above), the measured ellipsometric angles and
can be related to the optical properties of the thin film system by eqn 11.15. tan… †ej
ˆ  ˆ

r01p ‡ r12p eÿj2 1 ‡ r01s r12s eÿj2  1 ‡ r01p r12p eÿj2 r01s ‡ r12s eÿj2

11:15

It is assumed that the film is optically isotropic and homogeneous with a uniform refractive index N1, throughout its thickness, d. The ambient and substrate media are also considered semi-infinite, homogeneous and optically isotropic with refractive indices N0 and N2 respectively.12 Ellipsometry allows the measurement of two values, and
, which describe a change in the polarisation state caused by the sample. These two parameters are functions of the optical constants or the film thickness characteristics. For an ideal bulk substrate, the measured ellipsometric parameters, and
, can be directly inverted to give optical constants n and k. For more complicated structures, which include multiple layers, non-ideal interfaces, or gradients in the film optical properties, the analytical inversion of eqn 11.15 to extract the optical constants and the structural information is usually impossible. In such cases, advanced data analysis techniques are required to build a characteristic model of the material that is being measured, to extract simultaneously and uniquely the optical constants from the measured data. There are various numerical techniques and parametric models 12±33 for analysing ellipsometric data and the determination of the optical parameters of the sample. For a given ambient-film-substrate system, a single wavelength measurement of and
at one angle of incidence, 0 , provides enough information to

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determine only two real parameters of the system, assuming all the remaining parameters are known. Thus, for example, the complex refractive index of a film, n ÿ jk, can be determined if the film thickness, d, and the refractive index of the ambient and substrate are known. When the number of unknown parameters of the system exceeds two, as in most practical cases, (for example in a case where the film complex refractive index, n ÿ jk, and its thickness, d, are to be determined), additional data are required. There are a number of ways of increasing the number of independent ellipsometer readings12 including: · assuming the refractive index of the film stays unchanged, multiple measurement can be taken with varying of the thickness of the film · taking readings on a single film for different media or substrates, provided the film properties remain unchanged · recording independent data readings for a single film at various angles of incidence · making multiple wavelength measurements on the same film at a fixed angle of incidence. In practice, multiple angle of incidence (MAI) ellipsometry and multiple wavelength ellipsometry (MAE) are commonly used due to the simplicity and non-destructive nature of the measurement technique.34-37

11.2 History of ellipsometry and polarisation control 11.2.1 Instrumentation and optical elements Ellipsometry has been applied across a wide spectrum of fields in physics, chemistry, materials science and biomedical studies and optical and electronic engineering, for example. The steady growth of interest and wider application areas has led to the evolution of ellipsometry as a technique, having passed through a number of development stages. As a result, various laboratory, as well as commercial, ellipsometric configurations have been developed over a number of years. Common to most ellipsometers is a configuration consisting of two arms, one for generating an appropriate polarisation state of light to interact with a sample under measurement and another for collecting the reflected optical signal for detection and analysis. Usually the optical instrumentation comprises components such as polarisers, analysers and compensators, along with a monochromatic light source and a detection system.

11.2.2 The null ellipsometer The null ellipsometer3,6 is one of the earliest ellipsometric configurations to have been reported. The operational principle relies on manually or automatically turning a compensator and analyser axis so that the light reflected

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11.4 Schematic diagram of a null ellipsometer. Here L is a light source; P, a polariser; C, a compensator; S, a sample; A, an analyser; PD, a photo-detector.

from the sample is null at the detection point. The angular orientations of the polarising components are then reduced to yield the values of and
, by using a simple relationship.3 There are a number of null ellipsometric configurations, the most common of which, in principle, is the polariser-compensator-sampleanalyser (PCSA) arrangement,3,6 shown in Fig. 11.4. The compensator or quarter wave plate (QWP) is fixed at 45ë to the plane of incidence; the polariser and analyser are rotated until the null is found. It should be noted that near the null point, the change in intensity as a function of the azimuth angle is very small. In addition, near the null point, the detected light level is usually very low, indeed often reaching a level where the detector noise becomes a significant problem. To overcome this, a null may be obtained by using two angular positions on either side of the minimum light detection and then taking the average value. Such a manual null ellipsometer relies on careful calibration of the azimuth scales of the polarising components and adjustment of the optical components to reach the null point. For applications that require faster measurement speeds, or in cases where measurements may be repeated several times, the manual null ellipsometer is an impractical device, and an automated measurement becomes essential.

11.2.3 Automatic null instruments The principle of operation of these instruments is the same as for the manual null ellipsometer described earlier; however, in this case a servo-system instead of a manual adjustment is employed to achieve the null point. There are two kinds of automatic null ellipsometers, those that employ motors to drive the polarising components and electro-optic self-nulling ellipsometers with no moving parts. Early versions of motor driven null ellipsometers were those designed by Takasaki3 and Ord Wills.3 Here, two ADP (ammonium dihydrogen phosphate) modulators are used to rotate the polariser and the analyser in the PCSA ellipsometer. The response speed was rather slow, as five seconds was needed to

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reach a null point to 1ë accuracy. The AC voltage applied for the ADP cells was ~200 V, which is not attractive for use in a commercial instrument. Another limitation for Takasaki's servo null ellipsometer is the need for visual reading of the scales by the operator, once the null has been reached automatically. This problem was overcome by replacing the servomotors by stepper motors to drive the polariser and analyser.3 An automatic read-out of the polariser and analyser azimuthal settings was obtained by counting the pulses that drive the stepping motor to achieve the null. This system, which was reported, has 0.01ë resolution and one second nulling time. Winterbottom3 suggested a stationary self-nulling electro-optic ellipsometer, offering a high measurement speed and precision. The principle of operation of this device is to use Faraday cells to rotate the azimuth angles of the polariser and analyser magneto-optically, instead of using a manual or motor driven adjustment. The Faraday cells are driven by a variable DC current and an alternating modulation current. The component of the detector signal at the modulation frequency is fed back to control the DC current levels in the polariser and analyser Faraday cells3 until a null is reached. Ellipsometric parameters are then derived directly from the Faraday cell current supplies. Yamaguchi and Hasunuma3 developed a fast electro-optic self-nulling ellipsometer. Two KDP (potassium dihydrogen phosphate) crystals act as voltage controlled variable phase retardation plates, (a /2 retardation is induced by 75 kV at 546 nm), and by controlling the voltages applied to the KDP, any required polarisation state can be achieved. Automatic nulling is achieved by feedback control of the DC voltage levels applied to the KDP crystals. The voltage readout enables the values of and
to be obtained directly. The typical response time of an instrument of this type was of the order of one second.

11.2.4 Automatic photometric ellipsometer The second of the automatic ellipsometers is the photometric instrument. Such a device uses the values of the variation of light intensity for measuring the ellipsometric parameters and
. The intensity variation can be achieved as a function of one or more of the following: the azimuth angle, the retardation or the angle of incidence. Systems of these type include the rotating analyser ellipsometer3 and the polarisation modulated ellipsometer. The rotating analyser ellipsometer In 1975, Aspen and Stude34 introduced the rotating analyser ellipsometer (RAE). In this system, the analyser is synchronously rotating at frequencies ~100 Hz with its rotation axis being set along the direction of the beam propagation while the polariser and compensator remain fixed with defined azimuthal orientations.

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Plane polarised light reflected from a sample passes through a rotating analyser; the detected signal intensity shows a sinusoidal function from which and
are determined.18 The operation of the RAE is now well established and, in fact, different versions of the RAE and its counterpart, the rotating polariser ellipsometer (RPE) configurations,34±41 are commonly employed in commercial systems. Polarisation modulated ellipsometry The nulling techniques described thus far require several minutes for measurement. Such a long response time is undesirable in certain applications, such as a study of dynamic interfacial properties where rapid surface contamination is intolerable or when a fast process such as a chemical reaction or film deposition needs to be monitored.42±48 Jasperson et al.49 have introduced polarisation-modulated ellipsometry (PME) as an improved system to replace the null ellipsometric techniques that existed prior to that time. The principle of operation is that the state of polarisation of the light beam is modulated by introducing a controlled phase retardation so that the information carried on the reflected light from the optical sample under measurement is retrievable by analysing the resulting time-varying photo-electric current that is detected.3 Several possible configurations exist, the choice depending on the application of the modulation. A design proposed by Jasperson et al.49 is the polariser modulator sample analyser (PMSA) arrangement. In this arrangement, a fused quartz block is electronically driven, by an AC signal, through a piezoelectric crystal cemented to one end of the block. The oscillating strain induces a modulated birefringence (the dual refractive index axis). Thus the quartz block acts as a linear retarder with a time-varying retardation. The amplitude of the retardation is proportional to the applied varying voltage. In the PME design, all the optical components remain stationary. This freedom of mechanical movement is a distinct advantage and allows high-speed measurement. Since the work of Jasperson et al.,49 several researchers have worked on the PME approach and more attention has been given to this technique, due to the obvious advantages it offers for time resolved measurement applications. Shamir et al.50 have proposed a polarisation-modulated ellipsometer, which was free from mechanically moving parts. An acoustic modulator and a waveplate are used to generate the rotating plane polarised (RPP) beam. Singher et al.51 introduced an ellipsometric set up using a commercially available stabilised Zeeman laser, (SZL). An SZL emits a beam composed of two circularly polarised components that differ in frequency. This generates an RPP beam at their difference frequency, which is ideal for ellipsometric measurements but is expensive. With the advent of polarisation maintaining fibre, Yoshino and Kurosawa (1984)52 introduced the first ellipsometer based on a high birefringence (HiBi) fibre as a polarisation modulator. Work by the authors, discussed later, has

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demonstrated that an all-fibre ellipsometer could be used to measure the optical constants of the materials and operate as a surface sensor.

11.2.5 Spectroscopic ellipsometry Several single wavelength ellipsometric measurement techniques have been available for many years, and with on-going developments in optical sources and components, multi-wavelength or spectroscopic ellipsometry54±63 has become a powerful alternative for practical applications to characterise fully the optical properties of surfaces and thin films over a broad range of applications, by extracting the refractive index and film thickness, which are functions of the wavelength of light used. For example, several applications require ellipsometric measurements in the UV or IR region while others are best made in the visible part of the spectrum. Spectroscopic ellipsometry offers greater flexibility that can be used to meet a range of application requirements. In this method, the ellipsometric parameters, and
are measured at a given wavelength and the measurement is repeated for each wavelength across the spectral range of interest, thus determining n and k as a function of wavelength. These optical parameters are then matched to the output from appropriate computer models, to determine the structure and composition of the sample.

11.3 Fibre based polarisation modulated ellipsometry So far, the major progress and the direction of the research carried out in the area of ellipsometry has been directed to realising an accurate, flexible and compact measuring tool, which is capable of undertaking fast time resolved measurements. Polarisation modulated ellipsometry promises such features and has been the subject of research by several groups.49±53 With the recent developments in the telecoms industry and the advance of fibre optic technology, high birefringence (HiBi) fibre has been made more readily available and has been applied in developing an all fibre ellipsometer.52 Thus the combination of polarisation modulation and fibre-based ellipsometry offers the possibility of better and more affordable ellipsometers for a range of applications. The conventional ellipsometer consists of familiar polarising optical components that require precise alignment. Such a system is often physically large and bulky and thus is limited to use in the laboratory environment. Automation is usually achieved by mechanically moving one of the optical components for modulation, which limits the operational speed. This is undesirable for many applications where dynamic processes are to be monitored. A compact system would broaden the application areas of ellipsometry as a sensing tool or as a field instrument for environmental applications, potentially at lower cost than that of the present generation of commercial ellipsometers. An all-fibre ellipsometer offers a more compact and flexible system, immune to

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electromagnetic noise and enables remote sensing in hazardous environments. Such a system is free of mechanically moving parts, offering a potential for high-speed measurements, which does not rely on careful azimuthal alignment.

11.3.1 Polarisation state control and polarisation modulation There are a number of sensor schemes in which polarisation azimuth control and measurement forms a basis for the optical instrumentation used. These include ellipsometry and polarimetry, in which a well-defined polarisation-modulated beam is generated, such that upon the interaction of the beam with the sample, the changes in polarisation can be measured. These changes in polarisation are then used to characterise the sample. To date, several techniques have been proposed for producing a linearly polarised beam of controllable azimuth. Early work in the area had mainly concentrated on free space instrumentation using bulky optical components such as beam splitters, polarisers and birefringent crystals, for example. With the advance of fibre optic technology, a range of techniques has evolved for the generation of a modulated polarisation state.53,64,65

11.3.2 Polarisation maintaining behaviour in single mode fibres Conventional single mode optical fibres do not normally preserve the polarisation state of the light propagating along the fibre. In many applications in which the polarisation state of the light is irrelevant, this is of little significance. However, in polarimetric or interferometric sensor applications, ellipsometry and in coherent communication systems, the polarisation state of the light injected into the fibre should be allowed to propagate unchanged. This requires a true single mode fibre, which can support a single polarisation state along the entire length. Such a fibre would, in principle, have a perfect circular symmetric core along its entire length. In practice, in commercial `single mode' (or monomode) fibres, imperfections such as asymmetrical stress and core ellipticity affect the polarisation of the propagating light, i.e., the light beams propagating along the two normal eigenmodes of the fibre experience different refractive indices or birefringence and thus they propagate with differing phase velocities. The polarisation state is also dependent on external forces such as bends, twists, transverse pressure or temperature changes of the fibre. Thus for a conventional monomode fibre or low birefringence fibre, the SOP of the propagating light is not preserved. Polarisation-based measurement applications with such fibres require the attaching of a polarisation controller at the output end of the fibre and they are usually restricted to high stability (e.g. submarine) environments.66 Early work on polarisation control was directed towards exploiting the birefringence of the fibre by subjecting it to some kind of external perturbation

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through various elasto-optic,67±71 electro-optic64 or magneto-optic65,70 index changes. By using one of these methods, the circular symmetry of the ideal fibre is broken and as a result a phase shift is introduced between the two light polarisation components along the eigenmodes of the fibre. With an appropriate value of the phase shift, a controlled polarisation state can be achieved.72 In the absence of single polarisation fibres, the phenomenon of birefringence has been exploited to develop a high birefringence (HiBi) fibre that can preserve the SOP where the intrinsic birefringence is increased deliberately by imposing some form of asymmetry in the cross-sectional fibre structure. The level of intrinsic birefringence achieved is far greater than that which can be produced by external perturbations, and thus the fibre is immune to the effects of the unstable environment.73 A single mode HiBi fibre is able to transmit two orthogonal polarisation modes or eigenmodes with slightly different velocities. Such a fibre will preserve the linear polarisations of the propagating light when launched in one of the eigenmodes. With the introduction of HiBi fibres into fibre-optic technology, optical sensor design and polarisation control techniques rapidly enabled the replacement of precision bulky mechanical systems by fibreoptic equivalents providing more compact and cheaper systems. A higher speed of modulation was also achieved. Tatam et al. (1986)53 described a range of techniques in which polarisation control could be achieved, based on fibre-optic or acousto-optic modulators. Perhaps the most attractive scheme for polarisation control proposed by the above authors is the single birefringent fibre system with a piezoelectric transducer (PZT) modulator. A schematic diagram is shown in Fig. 11.5. In this scheme, light from a linearly polarised argon-ion laser was launched in to a HiBi fibre with its polarisation angle at 45ë to the fibre eigenaxes, to excite both the fibre orthogonal modes equally. At the output of the fibre, a quarter wave plate (QWP) oriented at 45ë to the fibre eigenaxis was used to return the beam to a linearly polarised state. Polarisation modulation was achieved by applying a longitudinal strain, modulating the modal phase retardance of the HiBi fibre with a saw-tooth waveform driving a PZT. This technique provides better polarisation control and is free of bulky mechanical components, since the modulation is achieved without bending, twisting or subjecting the fibre to either transverse pressure or electric or magnetic fields.

11.5 Polarisation state azimuth control using a single highly birefringent monomode fibre. Here L is an Ar‡ laser; PZT, a piezo electric transducer; F, HiBi fibre and /4 QWP, quarter wave plate.

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If the eigenaxes of a HiBi fibre are equally populated with linearly polarised light, ideally there should be a constant phase relationship between the two orthogonal plane polarisations introduced by the fibre; however, extrinsically induced birefringence effects, due to environmental effects, such as temperature variations, introduce a random phase variation. Johnson et al.74 have proposed a feedback controlled, HiBi fibre polarisation controller and their work has highlighted environmental effects on HiBi fibre-based systems and proposed phase variation compensation techniques. A number of schemes thus have been proposed for phase stabilisation in fibre optic systems.75±77

11.4 A high birefringence fibre polarisation modulation ellipsometry Work has been conducted by the authors on a HiBi fibre polarisation modulation ellipsometer.78±81 A much simpler scheme, with fewer optical components, is described, in which linearly polarised light is launched into a single mode HiBi fibre with its eigenaxes aligned at 45ë to the polarisation of the input light, to excite the two normal eigenmodes equally. The configuration used is shown in Fig. 11.6. Unlike the use of the previously described mechanically moving ellipsometers, the polarisation modulation is achieved by stretching a single mode HiBi fibre longitudinally,66 responding to a sinusoidal signal generator, so that the polarisation state of the incident beam is modulated to produce a phase generated carrier signal with a sinusoidal phase change, which can be represented as: …t† ˆ A sin …!t†

11:16

where A is the relative phase amplitude as a function of the peak to peak voltage of the driver signal, …t† is proportional to the birefringence change phase modulation and ! is the frequency of the driver signal. There is also a phase bias, B , associated with the static fibre length.82 Thus the overall phase relation, …t†, may be represented as: …t† ˆ B ‡ A sin …!t†

11:17

8

A Jones matrix analysis of the PMA (polariser-modulator-analyser) arrangement shows that the detected signal is of the form:

11.6 The PMSA experimental configuration. Here L is the HeNe source; P45 and A45, a polariser and analyser at 45ë with respect to the HiBi fibre eigen axis; PMU, a polarisation modulation unit and PD, a photodetector.

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11:18

where I…t† is a time variant intensity. The signal can be expanded into a series of Bessel functions43,44 such that, I…t† ˆ I0 ‰1 ‡ cos …B †J0 …A †Š‡ hX i 2I0 cos …B †J2m …A †cos …2!mt† ‡ hX i 2I0 sin …B †J2m‡1 …A † sin …2m ‡ 1†!t

11:19

where Jm …A † are Bessel functions of the first kind with order m and argument A . The above complex harmonic expression can be simplified by controlling the phase parameters, A and B . The phase bias, B , can be varied between zero and =2 by applying a controlled strain to the fibre so that eqn 11.19 could be simplified to contain either the odd or even frequency components along with a DC level.82 Further, the necessary simplification is also achieved by setting the amplitude of the modulating signal so that J0 …A † ˆ 0. At this point the amplitude of the modulating signal corresponds to a Bessel angle of 2.4048ë and J1 …A † and J1 …B † are within 12% of their peak values.49,84 Precision of less than 0.1% is easily achievable in this approach. A calibration procedure was followed to set J0 …A † ˆ 0, such that when the odd or the even harmonic frequency components were selected through varying the phase, B , between zero and =2, the amplitude of the shaker signal is set so that the DC level in either case must be the same. This means that the DC contribution due to J0 …A † in the signal is cancelled. An alternative procedure involves rotating the analyser at the output of the fibre until the modulation fringes disappear and the signal is reduced to DC, i.e., I…t† ˆ I0 . If the analyser is then set to 45ë and the amplitude of the shaker signal is adjusted until the same DC level as above is achieved, J0 …A † is then set to zero. At this point, rotating the analyser should not change the DC level in the signal. Once J0 …A † ˆ 0, by setting B , to either =2 or zero, eqn 11.19 can be simplified to: I…t† ˆ I0 ‰1 ‡ 2J1 …A † sin …!t† ‡ . . .Š

11:20

I…t† ˆ I0 ‰1 ‡ 2J2 …A † cos …2!t† ‡ . . .Š

11:21

or From these simplified expressions, J1 …A † and J1 …B † can be calibrated from the appropriate frequency component by controlling the phase bias, B , such that: R!cal ˆ

I! ˆ 2J1 …A † Idc

11:22

R2!cal ˆ

I2! ˆ 2J2 …A † Idc

11:23

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where I! , I2! and Idc are the intensities of the first two dominant frequencies and a DC intensity respectively, R! and R2! are ratios of the intensities of the first and second harmonic components and the DC component. For measurement applications, the calibrated signal must be stable at a certain frequency and amplitude.75,76,85 The behaviour of the HiBi fibre is such that the phase bias, B , responds to environmental perturbations53 which tends to vary the birefringence randomly. This phenomenon is apparent in the case when the polarisation axes of the fibre are both populated with normal polarisation light components, as is the case in the HiBi polarisation modulation fibre ellipsometry described here. The result is that the phase variation becomes random and thus the amplitude of the odd and even frequency components, as well as the DC level, fluctuates randomly. The basis of the ellipsometric scheme described so far is that the modulation unit permits some level of control of the state of polarisation of the output beam. In fact the fibre length extension, L(t), due to the modulation unit, is reflected in the phase shift introduced between the light travelling on the fast and the slow axes. This is due to the intrinsic properties of the highly birefringent fibre, specially manufactured to have a high degree of linear birefringence. In this configuration, it is realistically assumed that the phase modulating effect induced by the shaker is predominant with respect to the other uncontrolled environmental parameters, the effects of which can be neglected. Nevertheless, as was found experimentally, these unwanted factors influence the quality of the modulation of polarisation and as such the overall performance of the system. It is observed that the major extra modulation introduced in the system is due to the temperature variation and the air movement around the fibre, and in fact this effect was found to decrease by shielding the fibre in a small enclosed environment. Typically the temperature (T) response is 100 rad/ëC for a HiBi fibre such as was used experimentally (York 600HB).53 These variations affect the amplitude and shape of the harmonic content of the spectrum, but they are more evident in the first few harmonic frequencies of the spectrum, Idc , I! , I2! , due to the slowly varying nature of the temperature and other environmental effects. It has been proposed by the authors that the apparent problem of signal stability can be overcome by adding an extra controlled strain to the fibre in order to compensate the unwanted effect of  …T; t† by means of a feedback network. The extra phase shift introduced is such that the time varying phase shift is cancelled. A feedback control is thus employed to set the phase bias, B , at zero so that only the even frequencies and DC components are selected. A correction signal is generated by locking86,87 the free running interferometric signal to the driver signal. This phase sensitive signal is then fed back into an electro-dynamic shaker attached to the fibre section under tension, to cancel the effect of the unwanted external modulation effects on the fibre. Although an electronic feedback system has been effectively applied in the laboratory configuration, careful system design should be employed by making the fibre

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shorter and placing it in a thermally stable jacket. Further improvement could be achieved such as with the use of temperature-independent birefringent fibres proposed by Wong et al.,88 to reduce thermal drifts.

11.4.1 Measurement technique A sample under measurement, with its plane of incidence aligned to the fibre eigenaxis, is placed between the fibre and an analyser, which is oriented at 45ë to the same plane, as shown in Fig. 11.7. After a simple calibration procedure and feedback control to produce a DC and the even frequency harmonics, Jones matrix manipulation of the system components and the sample produces a reflected signal through a linear analyser with a detected intensity of the form: I…t† ˆ I0 rs2 ‰1 ‡ tan2 … † ‡ 2 tan … † cos……t† ‡
†Š

11:24

where I…t† is time varying intensity, I0 is the DC component and rs is the reflectance of the sample. Taking the first two dominant frequency components, eqn 11.24 can be expanded as: I…t† ˆ I0 rs2 ‰1 ‡ tan2 … †0 ÿ 4J1 …A † tan … † cos …!t† sin …
†‡ 4J2 …A † tan … † cos …
† sin …2!t†Š

11:25

The variations in the source intensity and transmittance of optical system can be rendered inconsequential by normalizing the component intensities with respect to the DC component. R! ˆ

I! ˆ 2J1 …A † sin …2 † sin …
† Idc

11:26

R2! ˆ

I2! ˆ 2J2 …A † sin …2 † cos …
† Idc

11:27

11.7 Experimental configuration of a HiBi fibre based ellipsometer with no quarter wave plate. Here P45ë and A45ë are polarisers at 45ë to the fibre eigen axes; PMU, a polarisation modulation unit; S, a sample and PD, a photodiode.

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where R! and R2! are the intensity ratios. The Bessel functions in the above relations can be eliminated by normalising these ratios against the calibration ratios of eqns 11.22 and 11.23. Thus: R! ˆ sin …2 † sin …
† R!cal

11:28

R2! ˆ sin …2 † cos …
† R2!cal

11:29

The above relations allow the determination of both and
from the two frequency components by taking a ratio of intensities. The simplicity of this signal processing technique is an advantage in real time, rapid process monitoring applications where faster, time-resolved measurements are important. and
measurements on a two phase system In demonstrating the approach discussed, ellipsometric measurements were carried out on glass and silicon substrates by the authors as a means of checking the accuracy of the system. The results from a glass sample (of refractive index 1.51) showed that more accurate readings are taken at higher angles of incidence, and this is in agreement with the observation that the system accuracy is much better for higher reflective samples. In particular, is more sensitive at higher incidence angles, due to the higher sensitivity of the sin(2 ) curve away from higher values. Measurements of and
for a silicon sample of refractive index 3.85ÿi0.018 at a range of incidence angles is shown in Figs 11.8 and 11.9.

11.8 Ellipsometric angles, and
as functions of incidence angles for reflection at an air/glass interface.

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11.9 Ellipsometric angles, and
as functions of incidence angles for reflection at an air/silicon interface.

The smooth curves are the theoretical values and the scatter curves are the experimental measurements. and
measurements on a three phase system Further thin film measurements were made on sol-gel films, carried out using the ellipsometric configuration shown in Fig. 11.6. Sol-gel films are important in several fibre optic sensor applications as they can be applied directly to fibres or substrates.89 Three thin sol-gel films of thickness of around 100 nm, 200 nm and 400 nm were used in this experiment, having first been analysed using a conventional laboratory ellipsometer. Similarly, and
were measured for each sample and these measured values were compared to values obtained by computation of the inversion process.90 Numerical inversion was carried out based on the assumption that the ambient refractive index of the air is 1 and of the silicon substrate, 3.858-i0.018, with the sol-gel samples having a refractive index of 1.58. The scatter curves in Figs 11.10 and 11.11 show the measurement results and the smooth curve shows the calculated and
values. The results reveal that may be measured to within an error of 1ë to 2ë while
has an error margin of 5ë for films of thickness 100 nm and 400 nm. It is interesting to note that the sample of thickness 200 nm showed a higher error margin as shown in Figs 11.9 and 11.10. This is due to the fact that this particular sample had been exposed to the laboratory environment for several weeks, and literally had reacted to the atmosphere as a result, while the other two were kept in a closed environment.

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11.10 measurements on three sol-gel films on silica substrate with various thicknesses at various angles of incidence.

11.11
measurements of three sol-gel films on silica substrate with various thickness at various angles of incidence.

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11.5 Future trends The advances in optical fibre systems continue to expand, in spite of the downturn in the optical fibre communications systems at the start of the 21st century. However, the technology developed remains key to advances in optical fibre sensor and measurement systems in general and thus to ellipsometry. The availability of specialist fibres has been a major benefit and possibilities exist for the use of exotic fibres with rare-earth dopants, for example, as light sources pumped by laser diodes to provide specialist optical fibre sources in compact allfibre systems. Multicore fibres are becoming more common, as well as a wider range of highly birefringent fibres. A major advantage for the optical instrumentation designer is the lower cost of many components and devices arising from the optical communications market ± higher value production and a more competitive market worldwide has made a significant difference here. Fibre Bragg gratings have become an important part of many fibre optical systems and, used in fibre reflection or filters, they have a significant role to play in this field as well. Overall, the move towards cheaper, lighter and more compact field deployable instrumentation enhances the potential of these future options more widely ± a trend seen very effectively, for example in the new generation of spectrometers available on the market. Coupled to advances in IT, the capabilities of virtual instrumentation software such as LabVIEW (from National Instruments), offer potential major advantages in specialist applications. Thus the field advances, based on sound optical principles, by taking full advantage of new technology as it becomes applicable and as new systems are developed.

11.6 Sources of further information As an established and versatile technology, ellipsometric measurement techniques have been widely applied both as research tools in various scientific environments and in manufacturing processes. As such there are several types of commercial ellipsometers available from different sources and an extensive body of further information is widely available, in addition to the literature listed in the references. These include commercial as well as academic and userinterest group websites. The goal here is not to provide every possible link to every ellipsometry based home page, but to include relevant and typical samples of what is available. J. A. Woollam Co http://www.jawoollam.com/ A leading supplier of spectroscopic ellipsometry, providing ellipsometers and associated data analysis software. There is an online tutorial on ellipsometry for

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the beginner in the subject. This is an excellent source of information about ellipsometry in general and spectroscopic ellipsometers in particular. Jobin Yvon Ltd. http://www.jobinyvon.co.uk The scientific components and instruments manufactured by the Jobin Yvon Group include single wavelength automatic ellipsometers and phase modulated spectroscopic ellipsometers with up to 50 kHz modulation frequency and fast data acquisition for dynamic studies of the optical properties of samples and liquid surface measurements, as well thin film characterisation. Gaertner Scientific Corporation http://www.gaertnerscientific.com Gaertner Scientific corporation offers laser ellipsometers that meet varied user requirements in the semiconductor, disc, magnetic head, flat panel and other thin film industries for non-contact thickness and refractive index measurements of thin transparent and semi-transparent films to 0.1 nm or better. This instrument uses no moving parts and no modulators to determine, quickly and accurately, the four Stokes parameters of the beam and thus its polarisation state from which the optical properties may be determined. Beaglehole Instruments http://www.beaglehole.com Beaglehole instruments provide a range of single wavelength and spectroscopic ellipsometers based on birefringence modulation for various applications including the semiconductor industries, optoelectronics and biotechnology. Rudolph Technologies, Inc. http://www.rudolphtech.com/products/index.html Rudolph Technologies provide process control metrology systems including microprocessor-controlled ellipsometric systems for transparent film metrology for use in the semiconductor industry. There are also other sources of information on the theory and applications of ellipsometry from several research and interest groups. Some of these can be obtained from the links below. http://www.uta.edu/optics/research/ellipsometry/ellipsometry.htm http://www.nrel.gov/measurements/ellips.html

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http://www.mfa.kfki.hu/icse-3/pages/presentations.html http://aip.org/tip/INPHFA/vol-10/iss-2/p14.html http://www.ieee.org/organizations/pubs/newsletters/leos/oct00/spectro.htm

11.7 References 1. Shurcliff, W. A., Polarised light: Production and use, Harvard University press, Cambridge, Massachusetts, 1969. 2. Clark, D. and Grainger, J. F., Polarised light and optical measurement, Pergamon Press Ltd., 1971. 3. Azzam, R. M. A. and Bashara, N. M., Ellipsometry and polarised light, Amsterdam: North Holland, 1977. 4. Vasicek, A., Optics of thin films, North-Holland Publising Co. Amsterdam, 1960. 5. Hect, E. and Zajac, A., Optics, Addison-Wesley Publishing Company, Inc., 1979. 6. Tomkins, H. G., A User's Guide to Ellipsometry, Academic Press, Inc., 1993. 7. Bashara, N. M., Buckman, A. B. and Hall, A. C., Recent developments in ellipsometry, North Holland Publishing Company, Amsterdam, 1969. 8. Hauge, P. S., `Recent developments in instrumentation in ellipsometry', Surf. Sci. 1980 96 108 9. Gerrard, A. and Burch, J. M., Introduction to matrix methods in optics, John Wiley and Sons, 1975. 10. Longhurst, R. S., Geometrical and Physical Optics, London, Longman, 1973. 11. Grant, R. F., Introduction to Modern Optics, Holt, Rinehart and Winston, Inc., 1968. 12. McCrackin, F. L. and Colson, J. P., `Computational techniques for the use of the Exact drude equations in Reflection Problems', Symp. Proc. On the measurement of surfaces and thin films. Misc. Publication 256, p 61±81, 1964. 13. McCrackin, F. L., Elio Passaglia, Robert R. Stromberg and Harold L. Steinberg, `Measurement of the Thickness and Refractive Index of Very thin Films and Optical Properties of Surfaces', A. Physics and Chemistry, 1961 67A(4) 362±377. 14. Mojun Chang and Ursula J. Gibson, `Optical constant determination of thin films by random search method', Applied Optics, 1985 24(4) 504±507. 15. Manifacier, J. C., Gasiot, J. and Fillard, J. P., `A simple method for the determination of the optical constants n, k and the thickness of a weakly absorbing thin film', Journal of Physics E: Scientific Instruments, 1976 9 1002±1004. 16. Case, William E., `Algebraic method for extracting thin film optical parameters from spectrophotometer measurements', Applied Optics, 1983 22(12) 1832±1836. 17. Drolet, J., Russev, S. C., Boyanov, M. I. and Leblanc, R. M. (1994), `Polynomial inversion of the single transparent layer problem in ellipsometry', Journal of Optical Society of America, 1994 (a) 11(12) 3284±3291. 18. Lekner, J., `Inversion of reflection elleipsometric data', Applied Optics, 1994 33(22) 5159±5165. 19. Dorf, M. C. and Lekner, J., `Reflection and transmission ellipsometry of a uniform layer', Journal of Optical Society of America (a), 1987 4(11) 2096±2100. 20. Holtslag, A. H. M. and Scholte, P. M. L. O., `Optical measurement of the refractive index, layer thickness and volume changes of thin films', Applied Optics, 1989 28(23) 5095±5104. 21. Loescher, D. H., Detry, R. J. and Clauser, M. J., `Least-Square Analysis of the FilmSubstrate Problem in Ellpsometry', Journal of the Optical Society of America. 1971

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61(9) 1230±1236. 22. Jellison Jr, G. E., `Data analysis for spectroscopic ellipsometry', Thin Solid Films, 1993 234 416. 23. Smith, D. S. P., Law, B. M., Smock, M. and Landau D. P., `Numerical analysis of ellipsometric critical adsorption data', Phys. Rev. E., 1997 55 620. 24. Wooten, F., Optical Properties of Solids, Academic Press, New York, 1972. 25. Terry Jr, F., `A modified harmonic oscillator approximation scheme for the dielectric constants of AlxGa1-xAs', J. Appl. Phys., 1991 70 409. 26. Jellison Jr, G. E. and Modine, F. A., `Parameterisation of the optical functions of amorphous materials in the interband region', Appl. Phys. Lett., 1996 69 371. 27. Jellison Jr, G. E., `Spectroscopic ellipsometry data analysis: measured versus calculated quantities', Thin Solid Films, 1998 33 313±314. 28. Kim, C. C., Garland, J. W., Abad, H. and Raccah, P. M., `Modeling the optical dielectric function of semiconductors: extension of the critical-point parabolic-band approximation', Phys. Rev. B, 1992 45 11749. 29. Zollner, S., `Model dielectric function for native oxides on compound semiconductors', Appl. Phys. Lett. 1993 63 2523. 30. Leng, J., Opsal, J., Chu, H., Senko, M. and Aspnes, D. E., `Analytic representations of the dielectric functions of materials for device and structural modeling', Thin Solid Films, 1998 132 313±314. 31. Herzinger, C. M., Yao, H. Synder, P. G., Celii, F. G., Kao, Y.-C., Johs, B. and Woollam, J. A., `Determination of AlAs optical constants by variable angle spectroscopic ellipsometry and a multisample analysis', J. Appl. Phys., 1995 77. 32. McGahan, W. A., Johs, B. and Woollam, J. A., `Techniques for ellipsometric measurement of the thickness and optical constants of thin absorbing films', Thin Solid Films, 1993 234 443. 33. Kihara, T. and Yokomori, K., `Simultaneous measurement of refractive index and thickness of thin film by polarised reflectances', Applied Optics, 1990 29(34) 5069± 5073. 34. Aspens, D. E. and Studna, A. A., `High precision scanning ellipsometer', Applied Optics, 1975 14 220±228. 35. Blaine, J. B., Woollam, J. A., Herzinger, C. M., Hilfiker, J., Synowicki, R. and Bungay, C. L., `Overview of Variable Angle Spectroscopic Ellipsometry (VASE), Part II: Advanced Applications', Critical Reviews of Optical Science and Technology, Proceedings of a SPIE, 2000 CR72 29±57. 36. Herzinger, C. M., Johs, B., McGahan, W.A. and Paulson, W., `A multi-sample, multi-wavelength, multi-angle investigation of the interface layer between silicon and thermally grown silicon dioxide', Thin Solid Films, 1998 313 281±285. 37. Maynard, H. L., Layadi, N. and Lee, J. T. C., `Plasma etching of submicron devices: in situ monitoring and control by multi-wavelength ellipsometry', Thin Solid Films, 1998 313 398±405. 38. Collins, R.W., `Automatic rotating element ellipsometers: Calibration, Operation, and Realtime Applications', Rev. Sci. Instrum., 1990 61 2029. 39. de Nijs, J. M. M. and van Silfhout, A., `Systematic and random errors in rotatinganalyser ellipsometry', J. Opt. Soc. Am. A., 1988 5 773. 40. Johs, B., `Regression calibration method for rotating element ellipsometers', Thin Solid Fillms, 1993 234 395. 41. Kleim, R., Kuntzler, L. and El Ghemmaz, A., `Systematic errors in rotating

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compensator ellipsometry', J. Opt. Soc. Am. A, 1994 11 2550. 42. Beaglehole, D., `Ellipsometry Studies of Liquid Surfaces', Proceedings International Conference of Ellipsometry, J. de Physique, 1983 44 C10-147. 43. Beaglehole, D., `Ellipsometric study of the surface of simple liquids', Physica, 1980 100B 163±74. 44. Beaglehole, D., Pfohl, T. and Riegler H., `An ellipsometric study of the surface freezing of liquid alkanes', Chemical Physics Letters, 1996 260 82. 45. Smith, D. S. P. and Law, B. M., `Ellipsometric study of critical adsorption and measurement of universal surface integrals', Phys. Rev. E, 1996 54 2727. 46. Manning-Benson, S., Bain, C. D. and Darton, R. C., `Measurement of Dynamic Interfacial Properties in an Overflowing Cylinder by Ellipsometry', J. Colloid Interface Sci., 1997 189 109±116. 47. Beaglehole, D., Wilson, P. and DeVries, A. L., `Antifreeze Glycopeptide Adsorption on single crystal ice surfaces by ellipsometry', Biophysical Journal, 1993 64 1878. 48. Peres, M. A. and Teijelo, M. L., `Ellipsometric study of WO3 films dissolution in aqueous solutions', Thin Solid Films, 2003 499 138±146. 49. Jasperson, S. N. and Schnaterley, S. E. `An improved method for high reflectivity ellipsometry based on a new polarisation modulation technique', Review of Scientific Instruments, 1969 40(6) 761±767. 50. Shamir, J. and Klein, A., `Ellipsometry with rotating plane polarised light', Applied Optics, 1986 25 1476±1480. 51. Singher, L., Brunfeld, A. and Shamir, J, `Ellipsometry with a Stabilised Zeeman Laser', Applied Optics, 1990 29 2405±2408. 52. Yoshino, T. and Kurosawa, K., `All Fibre Ellipsometry', Applied Optics, 1984 23 1100±1102. 53. Tatam, R. P., Jones, J. D. C. and Jackson, D.A., `Optical polarisation state control schemes using fibre optics or Bragg cells', J. Physics E: Sci. Instrum., 1986 19 711± 717. 54. Lee, J., Rovira, P. I., An, I. and Collins, R. W., `Rotating compensator multichannel ellipsometry: applications for real time Stokes vector spectroscopy of thin film growth', Rev. Sci. Instrum., 1998 69, 1800±1810. 55. Wethkamp, T., Wilmers, K., Esser, N., Richter, W., Ambacher, O., Angerer, H., Jungk, G., Johnson, R.L. and Cardona, M., `Spectroscopic ellipsometry measurements of AlxGa1-xN in the energy range 3-25 eV', Thin Solid Films, 1998 313 745±750. 56. Opsal, J., Fanton, J., Chen, J., Leng, J., Wei, L., Uhrich, C., Senko, M., Zaiser, C. and Aspnes, D. E., `Broadband spectral operation of a rotating-compensator ellipsometer', Thin Solid Films, 1998 313 58±61. 57. Aspnes, D. E., `Spectroscopic Ellipsometry of Solids,' Chap 15, Optical Properties of Solids ± New Developments, ed. B. Seraphin, North Holland, 1976. 58. Tiwald, T. E., Thompson, D. W., Woollam, J. A., Paulson, W. and Hance, R., `Application of IR variable angle spectroscopic ellipsometry to the determination of free carrier concentration depth profiles', Thin Solid Films, 1998 313 661±666. 59. Collins, R. W., An, I., Fujiwara, H., Lee, J., Lu, Y., Koh, J. and. Rovira, P. I., `Advances in multichannel spectroscopic ellipsometry', Thin Solid Films, 1998 313 18±32. 60. Synowicki, R. A., `Spectroscopic ellipsometry characterization of indium tin oxide film microstructure and optical constants', Thin Solid Films, 1998 313 394±397.

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61. Jellison Jr, G. E., `Examination of thin SiO2 films on Si using spectroscopic polarisation modulation ellipsometry', J. Appl. Phys., 1991 69 7627. 62. Herzinger, C. M., Johs, B., McGahan, W. A., Woollam, J. A. and Paulson, W., `Ellipsometric determination of optical constants for silicon and thermally grown silicon dioxide via a multisample, multi-wavelength, multi-angle investigation', J. Appl. Phys., 1998 83 3323. 63. Luttmann, M., Stehle, J. L., Defranoux, C. and Piel, J. P., `High accuracy IR ellipsometer working with a Ge brewster angle reflection polariser and grid analyser', Thin Solid Films, 1998 313 631±641. 64. Tkasaki, H., `Photoelectric Measurement of Polarised light by means of an ADP Polarisation modulator, II Photoelectric Elliptic polarimeter', Journal of Optical Society of America, 1961 51, 463. 65. Jackson, D. A., Kersey, A. D., Akhavan, L. P. and Jones, J. D. C., `High frequency non-mechanical optical linear polarisation state rotator', J. Phys. E: Sci. Instrum., 1985 19 146. 66. Noda, J., Okamoto, K. and Sasaki, Y., `Polarisation maintaining fibres and their applications', Journal of Lightwave Technology, 1986 LT-4 1071±1089. 67. Ulrich, R., `Polarisation stabilization on single mode fibre', Applied Physics Letters, 1979 35 840±842. 68. Kubota, M., Ohara, T., Furuya, K. and Suematsu, Y., `Control on single mode optical fibres', Electronics Letters, 1980 16 573. 69. Lefevre, H. C., `Single Mode Fibre fractional wave devices and polarisation controllers', Electronics Letters, 1980 16 778±780. 70. Okoshi, T., Fukay, N. and Kikuchi, K., `New Polarisation State control device: rotatable fibre Cranks', Electronics Letters, 1985 21 895±896. 71. Matsumoto, T. and Kano, H., `Endless Rotatable Fractional wave device for Single mode Fibre Optics', Electronics Letters, 1986 22 78±79. 72. Rahsleigh, S. C., Burns, W. K., Moeller, R. P. and Ulrich, R., `Polarisation-holding in Birefringence single-mode fibres', Optics Letters, 1982 7 40±42. 73. Varnham, M. P., Payne, D. N., Barlow, A. J. and Birch, R. D., `Analytic solution for the birefringence produced by thermal stress in polarisation maintaining optical fibres', Journal of Lightwave Technology, 1983 LT-1(2) 332±339. 74. Johnson, M. and Pannell, C. N., `Remote State of polarisation control on Polarisation maintaining fibre', Optics Communications, 1992 90 32. 75. Corke, M., Jones, J. D. C., Kersey, A. D. and Jackson, D. A., `All single-mode fibre optic holographic system with active fringe stabilization', J. Phys. E: Sci. Instrum., 1985 18 185±186. 76. Jackson, D. A, Priest, R., Dandridge, A. and Tveten, A. B., `Elimination of drift in a single mode optical fibre interferometer using piezoelectrically stretched coiled fibre', Applied Optics, 1980 19(17) 2926±2929. 77. Shadaram, M., Medrano, J., Papert, S. A. and Berry, M. H., `Techniques for stabilising the phase of the reference signals in analogue fibre-optic links', Applied Optics, 1995 34(36) 8283±8288. 78. Chitaree, R., Weir, K., Palmer, A. W. and Grattan, K. T. V., `A Highly Birefringent Polarisation modulation scheme for ellipsometry: System Analysis and performance', Measurement Science and Technology, 1994 5, 1226±1232. 79. Gebremichael, Y. M., Palmer, A. W., Grattan, K. T. V. and Weir, K., `Dynamic Oil Film Thickness Measurements using a HiBi Fibre Optic based Polarisation

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Surfaces and interfaces for biomaterials Modulation Ellipsometer', Divisional Conference of the Institute of Physics, Brighton UK. Proceedings of Applied Optics, 1998 169±174. Gebremichael, Y. M., Tedeschi, G., Palmer, A. W., Grattan, K. T. V. and Weir, K., `Fibre Optic Ellipsometer in a Thin Film based Sensor System Application, International Conference on Applications of Photonic Technology, Ottawa, Canada. Proceedings of SPIE/ICAPT98, 427g, 1998. Gebremichael, Y. M., Palmer, A. W., Grattan, K. T. V. and Weir, K., `Characterisation of oil films on water surface using fibre optic ellipsometry', International Conference on Applications of Photonic Technology, Ottawa, Canada, Proceedings of SPIE/ICAPT98, 427g, 1998. Grattan, K. T. V. and Meggit, B. T., Optical Fibre Sensor Technology, Chapman & Hall, 1995. Schilling, D. L. and Taub, H., Principles of Communication Systems, McGraw-Hill Book Co., 1986. Jellison Jr, G. E. and Modine, F. A., `Two channel polarisation modulation ellipsometer', Applied Optics, 1990 29(7) 959±974. Shadaram, M., Medrano, J., Papert, S. A. and Berry, M. H., `Techniques for stabilizing the phase of the reference signals in analogue fibre-optic links', Applied Optics, 1995 34(36) 8283±8288. Meade, M. L., `Advances in Lock-in Amplifier', J. Phys. E: Sci. Instrum., 1982 15, 395±403. Clayton, G. and Newby, B., Operational amplifiers, Butterworth-Heinemann Ltd., 1992. Wong, D. and Poole, S., `Temperature Independent Birefringent Fibres', International Journal of Optoelectronics, 1993 8 179±186. Badini, G. E., Grattan, K. T. V., Tseung, A. C. C. and Palmer, A. W., `Sol-gel properties for fibre optic sensor applications', Optical Fibre Technology: Materials, Devices and Systems, 1996 2 378±386. McCrackin, F. L. and Colson, J. P., `Computational Techniques for the Use of the Exact Drude Equations in Reflection Problems', Symp. Proc. On Ellipsometry in the Measurement of Surfaces and Thin Films, 1969.

12

Neutron reflection J R L U , UMIST, UK

12.1 Introduction For many biological and biotechnological applications, the solid/aqueous interface (S/L) represents the most appropriate model for revealing relevant molecular processes. A number of modern physical methods are capable of accessing the solid/aqueous interface in an in-situ environment. Techniques such as surface plasmon resonance (SPR) and ellipsometry are well established and they have been extensively used to follow a diverse range of biomolecular interactions. The technical strengths concerning SPR and ellipsometry are elegantly exemplified in refs 1±3, and are further reviewed in Chapters 10 and 11. In comparison, neutron reflection (NR) is a relatively new technique which is also capable of detecting molecular events at different interfaces, including the S/L.4,5 Although its application in biointerfacial studies is rather limited, its potential should not be underestimated.6±8 NR has two distinct features. First, it uses a short wavelength neutron beam as in the case of X-ray reflection. The Ê and is comparable to molecular size, wavelength ranges typically from 1±10 A making the measurement inherently more sensitive to the molecular dimension. Second, unlike X-ray reflection, neutron reflectivity can be altered by varying the isotopic labelling between H and D. Through selective labelling to a given component of interfacial species, its distribution in an in-situ environment can be identified, resulting in an improved interfacial structural resolution. The feature relating to isotopic labelling has direct implications for the design of interfacial experiments and is unique to NR. This review will utilise a number of examples from our recent studies to demonstrate the technical capabilities of NR. Because lysozyme is a widely used model protein, its recent study by NR will be selected here to demonstrate the effect of surface chemistry on protein unfolding. In addition, the structural features revealed from proteins of different size and globular stability will also be described. It is appropriate to mention here that biomolecular events at the solid/solution interface can also be studied in the form of particulates, vesicles or emulsion dispersions. Techniques such as circular dichroism (CD), NMR, small angle

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X-ray scattering (SAXS) and small angle neutron scattering (SANS) have been applied to reveal aspects of molecular features from such systems.9±18 Characterisation of wet biointerfaces represents a major technical challenge, however, new developments have been emerging over the past few years. A very noticeable progress is the development of CD for revealing secondary structures for proteins adsorbed at the planar interface,19,20 the use of infra-red-visible sum frequency generation (SFG) vibration spectroscopy to reveal fine structural details from proteins adsorbed at the solid/solution interface.21 It is very appropriate to emphasise that biointerfacial characterisation often requires a combined use of a range of physical and biochemical arsenals. A recent article by Merett et al.22 has elegantly reviewed the complementary roles of other contemporary physical techniques in studying different biointerfaces.

12.2 Neutron reflection and deuterium labelling Figure 12.1 depicts the experimental setup of neutron reflection, broadly similar to laser or infra-red reflection. A batch of parallel incidence beams is impinged onto a flat surface, and is reflected specularly onto the exiting side. The intensities of the incoming and exiting beams are monitored and the ratio of the two is termed reflectivity. As can be seen from the schematic shown in Fig. 12.1, there is a clear path difference between the upper and lower beams caused by the different points of contact at the interface, resulting in constructive and destructive interference fringes. This is expected, as from the normal light propagation. Reflectivity is often plotted as a function of momentum transfer, Q, where Q ˆ 4 sin = (where  is the incidence angle and  is the neutron beam wavelength). Changes in the level and shape of the reflectivity curves reflect different combinations of constructive and destructive interference fringes that are manifested by different interfacial film thicknesses and composition. Theories applicable to light reflection are also valid for neutron reflection. Neutron reflectivity, R, measured in abstract space, is often analytically related to the structure and composition of interfacial layers via the optical matrix formula, as described by Heavens,23 and Born and Wolff.24 Software is available that allows the calculation of reflectivity based on an assumed interfacial layer structure (thickness and neutron refractive index (which is usually termed

12.1 Schematic representation of neutron reflection from a solid-solution interface.

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scattering length density, )). The calculated reflectivity is then compared with that measured and the process is iterated until a good fit is obtained. It is often found that there may be more than one structural profile that fits the measured reflectivity well. However, the range of possible models is usually narrowed in the light of physically meaningful choices for a given interfacial system. A distinct feature of neutron reflection is parallel measurements under different isotopic substitutions without altering interfacial chemistry. Thus, the determination of more than one reflectivity profile helps to reduce the uncertainty in unravelling the real structural profile across the interface. The principle of application of isotopic labelling can be better illustrated by protein adsorption at the air/water interface, bearing in mind the schematic experimental geometry shown in Fig. 12.1. To optimise the signal from the adsorbed protein layer, it is necessary to reduce the contribution of signal from bulk water. This is achieved by mixing some 8 vol% D2O in H2O, or in the molar ratio of 1 to 11. As the resultant water has a scattering length density of zero, it is called null reflecting water, NRW. If the adsorbed protein is assumed to follow a Gaussian distribution, R can be expressed as RQ2 ˆ hpp ˆ ÿ2 exp …ÿQ2 2p =8† 162 b2p

12:1

where hpp is called protein partial structure factor, bp denotes the protein scattering length, ÿ is the surface adsorbed amount or surface excess and p is the full width of the protein layer at the height of 1/e.5,6 ÿ is related to area per molecule A through Aˆ

MW ÿNa

12:2

where MW denotes the molecular weight of the protein and Na denotes Avogadro constant. The conversion of reflectivity into layer thickness and surface coverage is well exemplified in Fig. 12.2 where the surface adsorption of human serum albumin (HSA) under different pH is compared.25,26 It can be seen from Fig. 12.2 that at pH 3 and 7 the two measured reflectivity profiles are almost identical. However, at pH 5, close to the isoelectric point (IP) of the protein, the reflectivity profile is higher in level over the low Q range and the entire curve has a steeper slope. Changes in the level and slope of reflectivity indicate a different amount of adsorption and layer thickness. The continuous lines through the measured reflectivity profiles represent the best fits from eqn 12.1 at the three different pHs, with layer thickness and surface excess adjusted accordingly. The results confirm that HSA adsorption reaches its maximum at Ê and pH 5, and that the layer thickness also reaches the greatest, at p ˆ 40  2 A 2 Ê 2, Ê A ˆ 5050 A . This value of A is slightly under the limiting value of 40  140 A assuming the molecules adsorb sideways on. Away from the IP, both layer

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12.2 HSA adsorption at the air/NRW interface at pH 3 (‡), 5 () and 7 (
), with the HSA concentration fixed at 1 gdmÿ3. The continuous lines represent the uniform layer model fits with ÿ ˆ 1:5  0:3 mgmÿ2 , P ˆ 33  3 Ð at pH 3; ÿ ˆ 2:2  0:3 mgmÿ2 , P ˆ 40  3 Ð at pH 5 and ÿ ˆ 1:6  0:3 mgmÿ2 , P ˆ 34  3 Ð at pH 7.

thickness and surface excess go down, indicating a strong correlation between the amount of HSA adsorption and lateral electrostatic repulsion. The concurrent attainment of maximal adsorption and layer thickness around the IP is indicative of the high structural flexibility of HSA molecules. It is useful to note that under all conditions, the adsorbed HSA layers were represented by a uniform layer model, indicating that in spite of structural deformations, the molecules retain their globular framework. Structural unfolding usually leads to the polypeptide distributions along the surface normal direction. Such a structural feature cannot usually be described by a single uniform layer model, as will be described later. Similar experiments at the air/D2O interface have also been made. The results have led to useful discussions about the extent of protein layer mixing with water. It is, however, relevant to mention that because the reflectivity now contains contributions from the protein layer, the water and their interactive term, its relation to the interfacial structural profile becomes less obvious. Likewise, when measured at the solid/water interface, the analytical expression linking reflectivity to the respective components across the interface will be further complicated. Information about layer thickness and composition is often derived from model fitting based on the optical matrix programme.27 In the following we should give a number of examples to illustrate the level of information that can be revealed about interfacial layers adsorbed at the solid/ solution interface. We will start with small peptide adsorption, and then proceed to large protein molecules.

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12.3 Peptide interfacial assembly The self-assembly of short synthetic peptides into molecular nano-structures is attracting growing interest and has been widely reported recently.28±33 The design of many of the primary sequences has taken into consideration the mimicry of those native proteins that in turn promote secondary structural elements of proteins, i.e., -helices, -sheets, turns, resulting in tertiary molecular arrangements amenable to the design of predictable and functional nano-structures. A number of recent studies have demonstrated nano-structuring in bulk solution from a diverse range of designed peptides. Many of these have also shown structural responses to solution pH and temperature. In collaboration with Professor S. Zhang at MIT, USA, we have recently shown that a number of short designed peptides form interesting nano-structures at the solid/water interface.34 We show here the adsorption of a 15-mer peptide that was mutated from a specific protein-binding domain within a heteromeric transcriptional activator HAP2 identified from yeast Saccharomyces cerevisiae.35 HAP2 is a 265 amino acid subunit that contains some 20 core amino acid region highly conserved for subunit assembly. Previous work by Zhang et al. has demonstrated that a 15-mer synthetic peptide within this region, between positions 162 and 176 in HAP2, showed a typical -helical circular dichroism (CD) spectrum. The sequence for this truncated peptide is shown in Fig. 12.3. The overlapping YYY side chains in the form of a -helical conformation constitute a characteristic tyrosine strand (and hence the notation of YYY). This, together with the adjacent hydrophobic strand of alanine and leucine and the positively charged strand of two lysines (administering their affinity through the two butylene chains), contributes strongly to the formation of a stable -helical structure as revealed by CD. As the solid substrate studied in this work was SiO2 and was weakly negatively charged, the positively charge groups provide an essential driving force for interfacial adsorption. Mutation of YYY into WWW is likely to alter the delicate balance of helical stability of the peptides, resulting in different degrees of peptide adsorption at

12.3 The primary sequences of wild YYY15 and mutant WWW15.

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12.4 Reflectivity profiles measured from the adsorption of 0.1 wt% wild YYY15 () and mutated WWW15 (
) at the hydrophilic silicon oxide/D2O interface at pH 7. The continuous lines represent a uniform peptide layer of 13 Ð and A ˆ 450 Ð2 for YYY15 and an -helical bilayer for WWW15 offering the structural pairing with the packing equivalent to A ˆ 230 Ð2 for each pair of peptides.

the solid SiO2/water interface. We thus compared the adsorption of the two peptides at pH 7 and 0.1 wt% and the reflectivity profiles measured are shown in Fig. 12.4. The resulting data analysis shown as continuous lines indicates a Ê with an area per molecule (A) of 450 A Ê 2 from peptide monolayer of 13 A YYY15 adsorption, showing that the peptide adsorbs sideways-on. Ê thick and more In contrast, the mutant WWW15 formed a layer of some 25 A importantly, this peptide layer had to be modelled using a symmetrical three Ê each. The middle sublayer had a scattering length sublayer model of some 8 A ÿ6 Ê ÿ2 density of some 2:5  10 A , close to that of the pure peptide. The scattering Ê ÿ2 , indicating the length density for the two side sublayers was some 5  10ÿ6 A association of some 50% D2O into these regions. The observed interfacial packing indicates the formation of a sideways-on peptide layer with complementary structural engagement between neighbouring WWW side chains. This structural integration is consistent with the possible formation of -helical peptide conformations. The neutron reflection experiment thus revealed the formation of a wellpacked sideways-on peptide layer at pH 7 for mutant WWW15, while the native YYY15 showed rather loose layer packing under the same solution conditions. However, when solution pH is lowered to 5, no adsorption was detected from WWW15 either. The reduction of surface affinity is in accord with the expected

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increased peptide hydrophibicity associated with protonation of Lys and Arg groups although increase in the opposite charge density would favour the adsorption. This situation was in strong contrast to the retaining of peptide adsorption when the peptide layer pre-adsorbed at high pH was exposed to low pH solution, indicating the high extent of irreversibility. These structural features are highly relevant to the exploitation of peptide interfacial assembly in bionanotechnology.

12.4 Lysozyme adsorption: the effect of surface chemistry Lysozyme is a small but robust protein that has been extensively studied.36±38 It Ê 3 with an IP around pH 11. has globular dimensions of some 30  30  45 A Thus, over the normal pH range, lysozyme is positively charged. The measurement of its adsorption at the solid/solution interface by neutron reflection has provided useful information for direct comparison with a vast amount of data from other techniques.

12.4.1 At the hydrophilic SiO2/solution interface The neutron reflection study of lysozyme adsorption at the solid/solution interface was carried out using surfaces representing a different chemical nature. The first set of neutron reflection measurements was made at the bare silicon oxide/water interface over a wide range of lysozyme concentration and pH.39,40 The bare SiO2 surface is very hydrophilic with contact angle close to zero. The surface is weakly negatively charged, and between pH 3 and 8, the charge density on the SiO2 surface does not vary much with pH. Thus, over the normal pH range the positively charged lysozyme has a natural tendency to adsorb onto the negatively charged SiO2. Figure 12.5 shows the pH dependent adsorption between pH 4 and 7. It shows that as pH increases the adsorbed amount also increases. This trend is opposite to the intuitive expectation of the increasing attractive force between the negatively charged SiO2 surface and the positively charged lysozyme molecules when the solution pH is shifted toward acidic condition. However, as the solution pH is moved close to the protein IP, the lateral repulsion between protein molecules within the adsorbed layer decreases. The results may simply show that lateral repulsion is more important in determining the amount of adsorption. Associated with the change of the adsorbed amount is the structural conformation of the adsorbed protein layer with pH. The decreased surface coverage with pH is accompanied with an increased extent of end-on orientation, as a result of increased repulsion between the molecules inside the monolayer. At higher bulk concentrations, adsorption produces a sidewayson bilayer at pH 7, and the measurements suggest that there is less lysozyme in

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12.5 Schematic representation of lysozyme adsorption at the hydrophilic silicon oxide/solution interface with lysozyme concentration fixed at 1 gdmÿ3. The solution pH was shifted between 4 and 7 and reversible adsorption was observed.

the outer layer. Only monolayer adsorption occurs at pH 4 because of the strong electrostatic repulsion within the protein layer but the higher coverage induced by higher bulk lysozyme concentration leads to a higher fraction of molecules adopting the end-on orientation. Further experiments were carried out to examine the reversibility of lysozyme by cycling solution pH, e.g., from 4 to 7 and back to 4, or vice versa. It was found that the adsorbed amount and layer thickness were determined by the final solution pH, and were not affected by the history that was experienced by the surface. However, the reversibility with respect to pH does not necessarily mean that the structure of the protein is not affected by adsorption. Direct contact between protein and solid surface may lead to partial breakdown of fragments of -helix or -sheet, which may generate further contact with the surface. On the other hand, the native state of lysozyme in aqueous solution is highly ordered with most of the polypeptide backbone having little or no rotational freedom. Although structural rearrangement may occur upon adsorption, the internal coherence of the lysozyme should prevent it from unfolding into loose random structures on the surface.41,42 Upon adsorption, the thicknesses of adsorbed layers of globular proteins are therefore expected to be comparable with their dimensions in aqueous solution and measurements of the layer thickness can then be used to assess the orientation of the protein molecules on the surface and whether they are distorted to any extent. This has been supported by reflectivity measurements under different isotopic contrasts. For example, the fit of the set of reflectivity profiles at a lower Ê for the concentration of 0.03 gdmÿ3 to a single uniform layer model of 30 A

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protein layer, after allowance has been made for the exchange of labile hydrogen atoms, suggested that protein molecules were adsorbed sideways-on at this concentration. Similar measurements at different water contrasts at pH 4 also Ê at this pH, although the showed that the thickness of the layer was about 35 A amount of protein in the layer was less. The increase in the thickness of the layer suggested that the layer was slightly tilted. The layer thickening was consistent with increasing charge density within the protein layer as the pH was lowered.39,40

12.4.2 At the hydrophobic OTS/solution interface The results described above indicate that adsorption of globular proteins onto the hydrophilic silicon oxide/water interface does not lead to the breakdown of the globular framework. To examine the effect of surface hydrophobicity, we have also studied lysozyme adsorption onto silica surfaces modified by endanchoring a self-assembled monolayer of octadecyl trichlorosilane (OTS).43 The effect of pH on lysozyme adsorption at the hydrophobic solid/water interface was investigated in a similar way to that described above at the hydrophilic solid/water interface. We subsequently examined the reversibility of lysozyme adsorption against solution pH and found that the adsorption was entirely irreversible at the hydrophobed solid/solution interface. The irreversibility may arise from the strong hydrophobic interaction between the peptide fragments and the solid surface, resulting in the breakdown of the globular assembly. If the denatured protein completely unfolds, there will be regions of loose random structure on the surface. Depending on the extent of denaturation, hydrophobic fragments will be segregated at the OTS surface, allowing hydrophilic fragments to extend into the aqueous solution. Through more elaborate modelling, it is possible to reveal more detailed information about the state of the adsorbed lysozyme at the OTS/solution interface. It should, however, be remembered that the work on the adsorption of lysozyme at the hydrophilic solid/solution interface showed that the thickness of the adsorbed layers could be related to the known dimensions of the protein in solution and that this was taken to support the tertiary structure of the protein remaining largely intact on adsorption. However, it was found that for the hydrophobic surface no model of the interfacial layer where the tertiary structure of the lysozyme was retained could be made to fit the observed profiles. This observation, even without more detailed model analysis, further confirms that lysozyme is denatured at the hydrophobic OTS surface. Once the effort of trying to constrain the dimensions of the unperturbed globular protein into the model fitting is abolished, it was relatively straightforward to fit the reflection data to structural profiles characteristic of synthetic polymers. When a single uniform layer model was used, the shape of

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the reflectivities could be largely approximated by a protein layer with a Ê and a protein volume fraction of 0.85. The thickness of thickness of 12±15 A Ê is equivalent to the length of the side chains of two average amino some 12 A acids, suggesting that the adsorbed protein layer is in the form of peptide chains with the hydrophobic side chains adsorbed on the surface of OTS. The poor fits in the uniform layer model must be caused by the neglect of the hydrophilic side chains extending into aqueous solution. When a two-layer model was used with the outer layer having a lower volume fraction to account for the hydrophilic peptide tails, the fitting to the reflectivity profiles was improved. The results from the two-layer model show that although the adsorbed amount of lysozyme at pH 4 and 7 is the same, the second layer is thicker at pH 4. This variation of the diffuse layer with solution pH suggests that the tail region contains charged groups. As described previously, the net charge on a lysozyme molecule depends on solution pH, and the number of charges decreases from 10 at pH 4 to 8 at pH 7.41,44 Thus at pH 4, the charge density within the protein layer is higher and, once the protein is denatured, the greater freedom of movement will allow stronger repulsion to cause the layer to be more diffuse. At pH 7, this repulsion is reduced and the protein layer becomes denser. The structural changes within this lysozyme layer are quite different from those on the hydrophilic surface, and provide further evidence of protein denaturation. Although the two-layer model fits all of the reflectivity profiles in different pH conditions, it is a very coarse grained model and one would expect the actual segment distribution profiles to be smoother, although the intrinsic resolution of the neutron reflection experiment is not great enough to distinguish the twolayer model from a continuous distribution. The smoothened distribution, mimicking the model previously described for the adsorption of polyethylene oxide (PEO) at the air/water interface,45 presented a much better description for the loops or tail fragments in the diffuse region composed of either the hydrophobic fragments forced into the diffuse layer due to steric constraints, or peptide chains containing more polar or charged groups. The schematic representation of the peptide distributions at the hydrophobic OTS/solution interface is shown in Fig. 12.6. It should be noted that if polar and charged groups dominate the composition of the diffuse layer, the scattering length for the protein fragments in this region would be different from the average value. In the case of D2O the scattering length of the protein fragments in the diffuse layer will tend to be higher due to the exchange of a greater number of labile hydrogens; the current use of the average scattering length density from the polypeptide might then underestimate the amount of protein in the diffuse region. It is, however, difficult to estimate such errors because they depend on the exact content of the components in the diffuse layer, which is not known. However, since the fraction of protein in the diffuse layer is in all cases less than 0.2 of the total surface excess, the

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12.6 Schematic representation of lysozyme adsorption at the hydrophobic OTS/solution interface. The main features from the adsorbed layers are structural unfolding and irreversible adsorption.

uncertainty will not lead to serious errors in the values of the total surface composition. Adsorption of lysozyme onto hydrophobic surfaces has also been studied by other methods. Using ellipsometry,46 it was estimated that the adsorbed layers Ê. could be modelled by a uniform layer model with a thickness of around 100 A Such a large discrepancy is likely to be caused by the insensitivity of the elliptical signal to the protein layer that is heavily mixed with water. X-ray reflection has also been used to study protein adsorption, e.g., a recent study by Petrash et al.47 on the adsorption of HSA on a self-assembled monolayer of hexadecyl trichlorosilane on silica. In comparison with neutron reflection, a typical X-ray reflection can measure reflectivity up to 10ÿ8 with  extending out Ê ÿ1. This higher sensitivity renders X-ray reflection a greater to about 0.6 A structural resolution, but the main drawback for X-ray reflection at the moment is that such experiments can only be done ex situ and the subtle changes characteristic of proteins in an aqueous environment are lost.

12.4.3 At the C15OH/solution interface The hydrophilic silicon oxide and the hydrophobic OTS surface represented two extreme surface hydrophobicities. In a systematic effort to examine the effect of surface hydrophobicity, we have also studied the adsorption of lysozyme at the C15OH/water interface,48 where C15OH represents pentadecyl trichlorosilane with terminal hydroxyl groups (Cl3Si(CH2)15OH, abbreviated to C15OH). Its attachment onto a planar silicon oxide surface produced a stable self-assembled monolayer of C15OH with hydroxyl groups on the outer surface. The advancing contact angle (a ) for the C15OH surface was found to be around 52o, which was in the middle between the bare silicon oxide (c. 0ë) and OTS (c. 110ë). As already shown, the OTS surface caused a complete unfolding of globular proteins such as lysozyme, whilst the globular frameworks were largely retained

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when these protein molecules adsorbed on to the hydrophilic silicon oxide. The C15OH surface presented the intermediate hydrophobicity for direct comparison with the other two model surfaces in terms of protein adsorbed amount and structural unfolding. Neutron reflection was first carried out to characterise the quality of the C15OH layer at the solid/D2O interface. The result showed that the C15OH layer was well represented by a packing equivalent to liquid alkanol density with a Ê , indicating a small but negligible uniform layer density and a thickness of 16 A quantity of defects from water penetration. The thickness of the layer is shorter Ê , suggesting that the layer is on average than the fully extended length of 20 A tilted. It should be noted that in comparison with the attachment of organic monolayers on gold through thiol reaction (see the work of Prime et al.49), the coupling of trichlorosilane groups with the hydroxy groups on the surface of silicon oxide is less selective. The poor binding on silicon oxide has resulted in a less well packed organic layer, evidenced by the high advancing contact angle, as compared with some 10 to 30ë for the hydroxyl layers coated on gold. Furthermore, although neutron reflection offers a reliable measurement of the average volume fraction of the C15OH layer, it is insensitive to the possible presence of defects within the surface region. In contrast, other techniques such as IR can offer detailed information about the trans/gauche conformations of the underlying carbon chains, but are completely ineffective when such thin layers are grafted on silicon oxide. Lysozyme adsorption at the C15OH/D2O interface at pH 7 and 1 gdmÿ3 was Ê . The best fitted to a uniform layer distribution with the layer thickness of 38 A volume fraction of lysozyme within the layer is below 0.2, showing that the adsorbed layer is very loose. This layer thickness is between the two axial lengths of crystalline structure for lysozyme, indicating the formation of a tilted monolayer. That no density gradient is required to fit the data suggests that the adsorbed lysozyme also retains its globular framework. The sensitivity of the measured reflectivity to the thickness of the layer was tested by varying thickness away from the optimal value, while  was varied accordingly so that minimum deviation was caused between the calculated and the measured reflectivities. The large variation of the shape and level of the reflectivity with layer thickness strongly indicated that the measurement was sensitive to the Ê. thickness of the layer within the error of a few A Ê The formation of a tilted monolayer of 38 A at 1 gdmÿ3 and pH 7 is in contrast to the sideways-on adsorption under the same solution conditions at the SiO2/water interface. There are two main factors contributing to the observed difference. First, the presence of C15OH layer changes the level of van der Waals interaction between the surface and the protein layer.50 Second, the coating of the C15OH provides a spacing between SiO2 and the protein layer. As a result, the adsorbed lysozyme molecules are no longer in direct contact with the weakly negatively charged surface and the preferred conformation may

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result from the reduced electrostatic interaction between lysozyme and the coated surface. These results together with the concentration dependent adsorption show that in addition to conformational differences the amount of adsorption at the C15OH/solution interface is substantially more reduced than on the other two interfaces. The reversibility with respect to pH was examined similarly by fixing lysozyme concentration at 1 gdmÿ3. Each of these reflectivity profiles at the shifted pH was compared with that corresponding to the first exposure of the fresh surface under the same pH. Each reflectivity pair was identical at a fixed pH, which suggests that the adsorption is completely reversible. However, when the same study was carried out at a higher lysozyme concentration of 4 gdmÿ3, a large discrepancy between the two reflectivity profiles measured at pH 7 was found, suggesting that the adsorption is irreversible. Irreversible adsorption is likely to be caused by the pre-adsorbed lysozyme at pH 4. When the pH was increased to 7, the pre-adsorbed layer was largely retained, thus presenting a surface that was different from the hydroxyl groups. This pre-adsorbed lysozyme layer must have induced further lysozyme adsorption, resulting in a total amount greater than that obtained at the fresh C15OH-solution interface. The effect of surface hydrophobicity on protein adsorption has been rigorously examined by Prime et al.49,51 In their work the functionalised surfaces were formed by self-assembly onto Au and Ag through thiol reaction. Lysozyme adsorption on the surfaces with terminal hydroxyl groups, with ethoxylate units bearing terminal OH and OCH3 and with the mixtures of the ethoxylates and alkyl chains was investigated. Whitesides et al.49,51,52 found that a for self-assembled monolayers with terminal hydroxyl groups may vary over a wide range, but mostly below 30ë, as compared with some 50ë observed on silica, suggesting that a is affected by the packing of the coated monolayer and that the layers are overall better packed through thiol reaction than through silane coating. The effectiveness of the hydroxyl surfaces in prohibiting protein adsorption were found to be broadly comparable to those grafted with ethoxylate units. In addition, these authors found that layers coated on the surface of Ag give higher levels of trans conformation and hence lower contact angles, indicating better packing on the outer surface by oxygen atoms. The surface excess for lysozyme was, however, found to be much greater on the coated Ag surfaces, suggesting that surfaces with greater hydrophilicity produce higher surface excesses. A further piece of information that supports this trend of adsorption is our recent study of protein adsorption on poly(methylmethacrylate) (PMMA). The contact angle of PMMA is about 80ë and it was found that surface excesses at the PMMA/solution interface are also lower than the corresponding values on OTS and SiO2. It should, however, be noted that the correlation between the level of surface adsorption and surface hydrophobicity through contact angle is very empirical, because contact angle itself offers no information about the

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structure and chemical nature of the surface. An intermediate value of contact angle can be obtained through different means. For example, a contact angle of 50ë can be achieved by coating a loose layer of short chain alkanes, by forming a mixed layer of alkyl chains and hydrophilic groups. It is possible that adsorption varies with the nature of these surfaces, given that the surface contact angles and surface conditions remain the same.

12.4.4 At the phosphocholine (PC) monolayer/solution interface Extensive literature studies have now shown that surface coatings of thin polymer films containing phosphorylcholine groups are extremely effective at reducing non-specific protein deposition and rendering improved surface biocompatibility (e.g. see refs 54±56). Improved haemocompatibility associated with the protein repelling effect has also been observed when polymers were exposed to blood.57,58 Our study has been confined to the adsorption of pure proteins. This has allowed us to establish some straightforward relationships between the structure and composition of the adsorbed proteins and surface chemical nature, avoiding complication from blood at this early stage. It is relevant to note that self-assembly of phospholipid bilayers through physical attachment also shows high efficiency in reducing protein adsorption.57 However, the main shortcoming with the strategy of mimicking a labile phospholipid bilayer structure is the lack of stability stemming from the physical coating. Our PC monolayer was again self-assembled onto a silicon oxide surface through silane anchoring. The PC compounds used for surface coating were prepared in bulk solution, and the surface layers were formed via dip coating. A representative PC monomer we have prepared was APTMS-APC that was made by refluxing 3-aminopropyl trimethoxysilane (APTMS) with acryloyloxyethylphosphorylcholine (APC, synthesised using a procedure similar to the formation of PMMA) in isopropanol.58 Since the PC compound has a labile hydrogen on the secondary amine group, its dimer form was obtained by coupling two monomers with a bridging spacer such as a diisocyanate. The coupling was again carried out in isopropanol by refluxing the monomer and the spacer. The chemical structure of the APTMS-APC dimer is shown in Fig. 12.7. Clearly, this structure allows direct control of the surface coverage of the PC groups by the length of the spacer. It is appropriate to acknowledge that Hayward et al.59 had already explored the concept of the chemical grafting of PC monolayer onto silicon oxide some twenty years ago. But in their case, the attached alkyl chain bearing terminal PC groups was unstable as a result of the presence of an oxygen between carbon and silicon, the C-O-Si connection. The resulting alkyl silicate structure readily hydrolyses, with loss of the alkyl chains and their PC functionality. Our

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12.7 Molecular structure of a representative PC dimer used for surface grafting via silane chemistry.

approach is similar to that of Hayward et al. but we have sought direct chemical bonding between the carbon and silicon so that stable chemical grafting is obtained. In comparison with C-O-Si bonding, the formation of Si-O-Si connections at the end of the organic monolayer is much more stable. This part of the layer is identical to the underlying silica layer in chemical composition although their structures might be different. As in the case of protein adsorption onto a C15OH/solution interface, our NR study began with the characterisation of PC monolayer structure.60 The PC layer Ê containing 35% water. The high water content within the was found to be 18 A layer was likely to be caused by the bulky PC head groups. Protein adsorption was first characterised using lysozyme solutions with the lysozyme concentrations of 0.03, 1 and 4 gdmÿ3, all at pH 7 and in D2O. The main observation from these measurements was that the adsorbed lysozyme layer is thick but very diffuse. At the lowest lysozyme concentration of 0.03 gdmÿ3, the Ê thick and the volume fraction of protein is 0.08, correslayer is some 60 A ponding to a lysozyme surface excess of 0.8 mgmÿ2. At the highest Ê and ÿ ˆ 1:8 mgmÿ2, the fraction of the concentration of 4 gdmÿ3,  ˆ 68 A protein in the layer (fp) was 0.17. The feature of the formed thick and diffuse protein layers is remarkably similar to the observation found on the PC polymer surface,54±56 but is very different from the structural conformations observed on the bare oxide and OTS surfaces, as will be discussed later. The adsorbed lysozyme was found to be largely removable by rinsing the surface with pure buffer solution. Furthermore, the extent of lysozyme removal was greater than achieved on other surfaces as described above. Complete removal of adsorbed proteins was achieved by exposing the PC surfaces to a cationic surfactant, dodecyl trimethyl ammonium bromide (C12TAB). In comparison, only partial removal was obtained for lysozyme layers adsorbed on OTS and hydrophilic silicon oxide surfaces. These differences suggest that the proteins adsorbed on PC surfaces are loosely attached. We have also extended the assessment of PC monolayer surface by adsorption of BSA (bovine serum albumin) at a concentration around 1 gdmÿ3

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and at pH 5. The structural parameters obtained also showed the formation of a Ê thick but diffuse BSA layer distribution, with the layer thickness around 80 A ÿ2 with fp ˆ 0:08 and ÿ ˆ 1 mgm . These studies were then compared with fibrinogen adsorption in comparable solution concentrations at pH 5. Data analysis shows that the loosely adsorbed protein could also be approximated to a Ê with fp ˆ 0:08 and ÿ ˆ 0:5ÿ0:8 mgmÿ2. These diffuse layer of some 80 A results show that whilst the surface affinities of the proteins increase with their globular sizes, the extent of adsorption onto the PC monolayer surface does not show a positive correlation. In fact, the amount of adsorption shows a slightly decreasing trend with increasing protein size. This trend is opposite to that observed on other types of surfaces as will be discussed below.

12.5 Effect of size of globular proteins on their adsorption To compare the trends outlined above for the adsorption of different proteins onto the PC surface, an outline of their adsorption at the hydrophilic silicon oxide/solution interface is now given. These results together will allow direct comparison of structural features arising from the adsorption of these globular proteins onto two different interfaces. Following lysozyme adsorption as outlined previously, we have also investigated the adsorption of BSA and HSA at the hydrophilic silica/water interface.61,62 As the isoelectric point of BSA and HSA is 4.8, we found that the adsorbed BSA amount peaks at its IP, and that the surface excess decreases when the pH is moved away from the IP. The effect of pH was more intensified when the bulk concentration was increased. HSA showed a similar pattern of adsorption, except that HSA had a slightly lower surface excess. The difference is consistent with the observation at the air/water interface, confirming that BSA has a higher surface activity.25,26 The general pattern of the adsorption of globular proteins at the hydrophilic solid/water interface is remarkably similar to that described for their adsorption at the air/water interface. These are reflected in pH dependence with respect to the IP, and changes of surface excess with respect to bulk concentration. In most cases, the differences in surface excess at the two interfaces are within 1 mgmÿ2. These results show that as far as these two interfaces are concerned, the structure and composition of the protein layers are governed by electrostatic repulsion within the adsorbed layers. The interaction between the protein and the surface appears to be less significant. As in the case of lysozyme, the reversibility of albumin adsorption at the hydrophilic solid/water interface was also examined. Adsorption was found to be reversible with respect to pH when BSA concentration was less than 0.15 gdmÿ3, suggesting no denaturation.61,62 This observation was consistent with the successful fitting of measured reflectivity to a uniform layer model. As described in the case of adsorption onto an OTS surface, denaturation led to a

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more fragmented peptide distribution which required a heterogeneous density profile to describe the polypeptide distribution. Comparison of layer thickness with the dimensions of the ellipsoidal globular solution structure again indicated that the molecules adsorbed sideways-on. But the layer thickness was in many Ê , suggesting a different extent of cases well below the short axis length of 40 A structural deformation. An observed increase of layer thickness with bulk concentration indicated a reduced extent of structural distortion with surface excess. At a concentration of 2 gdmÿ3, adsorption was found to be irreversible to pH. The irreversibility was evidenced by the higher surface excess found when pH was increased from 5 to 7 compared with surface excess obtained from direct BSA adsorption onto the bare oxide surface at pH 7. As already explained, this difference could be caused by an insufficient energy to detach all the contacts simultaneously from the solid substrate. The adsorption of fibrinogen onto the SiO2/solution interface resulted in a similar adsorbed amount to that at the hydropbobic OTS/solution interface under comparable solution pH and concentrations. However, the density profile at the OTS/solution interface was found to be broadly similar to that described already for lysozyme adsorption at the same interface. In contrast, the structure of fibrinogen adsorbed at the SiO2/solution interface is represented by a threesublayer region, the inner layer close to the oxide surface containing mainly water, a middle sublayer containing mainly polypeptide and an outer diffuse layer containing mainly water again.63 The presence of the inner water was rather intriguing, but without introducing this sublayer, it was impossible to account for the shape of the neutron reflectivity profile as shown in Fig. 12.8.

12.6 Conclusion An important role of protein adsorption is its mediation of cell recognition, attachment and the subsequent proliferation. Understanding the effect of surface chemistry underlying this interactive process has direct implication on tissue engineering, biosensor and biomedical device development. Over the past few years, concerted effort in this area has been made by Whitesides and coworkers.64±67 In their studies, SAMs with terminal functional groups were formed on gold using well established thiol chemistry. Some 50 different functional groups were anchored onto surfaces, and their interactions with a range of representative proteins were screened. The extent of protein adsorption on different surface functionalities was then compared with the attachment and viability of representative mammalian cells and bacteria. Their endeavour thus marks the most comprehensive approach so far. Their studies strongly indicate that adsorbed protein layers can vary in adsorbed amount and structural conformation, and that influence on cell attachment can be both chemical and structural. Exact mechanistic processes at the molecular and cellular levels still remain complex. Our development of NR and other methods for direct

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12.8 Neutron reflectivity profiles measured at 0 (dashed line), 0.01 (ú) and 0.2 () gdmÿ3 of fibrinogen concentration, with fitted lines to the two higher concentrations shown as continuous lines. The broad interference fringes over the low Q range indicate the presence of diffusive protein layers. In addition, the model analysis revealed that it was essential to incorporate a water sublayer of some 10 Ð on the SiO2 surface, followed by a dense polypeptide sublayer of 15±20 Ð, and a further water rich diffuse outer sublayer of 20 Ð on the solution side.

measurements at the solid/solution interface will complement this effort by revealing the various in situ structural conformations. The structural details revealed from our limited experiments so far at different solid/solution interfaces have convincingly demonstrated the technical feasibility and future potential of applying these methods to various biointerface characterisations. Our NR studies, described above, have already revealed a number of interesting features about blood protein adsorption at the solid/solution interface. It is worthwhile to summarise them here. First, the amount of protein adsorption is reduced on surfaces with intermediate hydrophobicity. This is well demonstrated from the four model surfaces described in this chapter. Second, although adsorption is comparable between hydrophilic SiO2/solution and hydrophobic OTS/solution interfaces, the structural conformations of globular proteins are distinctly different. The OTS surface induces strong structural unfolding and denaturation, whilst protein layers adsorbed at the SiO2/solution interface are well represented by structural features matching the globular dimensions of protein framework. These observations are entirely consistent

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with the pH reversibility of adsorption at the SiO2/solution interface and the irreversibility at the OTS/solution interface. Third, the amount of adsorption at the PC monolayer/solution interface shows a decreasing trend with increasing size of globular protein, whilst the opposite trend is observed on all other interfaces involving SiO2, OTS, PMMA (poly(methyl methacrylate)) and PS (poly(styrene)). Furthermore, in addition to the reduced adsorbed amount, protein layers formed at the PC monolayer/solution interface tend to form a broad distribution with low protein volume fraction, indicating a low fraction of direct contacts with the substrate surface. This is entirely consistent with the high degree of reversibility of adsorption. These features may arise from the high extent of surface hydration and could be highly relevant to the distinct biocompatibility of these types of biomaterials. Direct measurement of in situ structural conformations of proteins at the solid/solution interface has been difficult. Many existing techniques find it either difficult to access this type of interface or lack structural sensitivity and resolution. The capability of NR in unravelling structural details inside the protein layer will stimulate further interest in its application to study proteins that are of more direct relevance to biomedical and biotechnological applications.

12.7 Acknowledgements The author wishes to acknowledge the support of this work by BBSRC, EPSRC and Biocompatibles UK Ltd. He would also like to thank the American Chemical Society (ACS) and the Royal Society of Chemistry (RSC) for permission to reproduce a number of figures in this chapter.

12.8 References 1. Green RJ, Frazier RA, Shakesheff KM, Davies MC, Roberts CJ, Tendler SJB, `Surface plasmon resonance analysis of dynamic biological interactions with biomaterials', Biomaterials, 2000 21(18) 1823±1835. 2. Ostuni E, Grzybowski BA, Mrksich M, Roberts CS, Whitesides GM, `Adsorption of proteins to hydrophobic sites on mixed self-assembled monolayers', Langmuir, 2003 19(5) 1861±1872. 3. Keddie JL, `Structural analysis of organic interfacial layers by ellipsometry', Curr Opinion Colloid Interface Sci 2001 6 (2) 102±110. 4. Hayter JB, Highfield RR, Pullman BJ, Thomas RK, Mcmullen AI, Penfold J, `Critical reflection of neutrons ± a new technique for investigating interfacial phenomena', J Chem Soc Faraday Trans I, 1981 77 1437±1448. 5. Lu JR, Lee EM, Thomas RK, `The analysis and interpretation of specular neutron and x-ray reflection', Acta Cryst, 1996 A52 11±41. 6. Lu JR, Thomas RK, `Neutron reflection from wet interfaces', J Chem Soc, Faraday Trans 1998 94 995±1018.

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7. Lu JR, Thomas RK, `The application of neutron and x-ray specular reflection to proteins at interfaces', chap. 18, pp. 609±650, in Physical Chemistry of Biological Interfaces, eds, Baszkin A and Norde W, Dekker, New York, 2000. 8. Lu JR, `Neutron reflection study of globular protein adsorption at planar interfaces', Annu Rep Prog Chem C, 1999 95 3±45. 9. Baptista RP, Santos AM, Fedorov A, Martinho JMG, Pichot C, Elaissari A, Cabral JMS, Taipa MA, `Activity, conformation and dynamics of cutinase adsorbed on poly(methyl methacrylate) latex particles', J Biotech, 2003 102 (3) 241±249. 10. Wang SX, Sun YT, Sui SF, `Membrane-induced conformational change in human apolipoprotein H', Biochem J, 2000 348, 103±106. 11. Tian MH, Lee WK, Bothwell MK, McGuire J, `Structural stability effects on adsorption of bacteriophage T4 lysozyme to colloidal silica', J Colloid Interface Sci, 1998 200 (1) 146±154. 12. Kondo A, Fukuda H, `Effects of adsorption conditions on kinetics of protein adsorption and conformational changes at ultrafine silica particles', J Colloid Interface Sci, 1998 198 (1) 34±41. 13. Billsten P, Freskgard PO, Carlsson U, Jonsson BH, Elwing H, `Adsorption to silica nanoparticles of human carbonic anhydrase II and truncated forms induce a moltenglobule-like structure', FEBS Lett, 1997 402 (1) 67±72. 14. Smith LJ, Clark DC, `Measurement of the secondary structure of adsorbed protein by circular dichroism. 1. Measurements of the helix content of adsorbed melittin', Biochem Biophys Acta, 1992 1121 (1±2) 111±118. 15. Kondo A, Murakami F, Higashitani K, `Circular dichroism studies on conformational changes in protein molecules upon adsorption on ultrafine polystyrene particles', Biotech Bioeng, 1992 40 (8) 889±894. 16. Ishihara K, Iwasaki Y, Nakabayashi N, `Polymeric lipid nanosphere consisting of water soluble poly(2-methacryloyloxyethyl phosphorylcholine-co-n-butyl methacrylate)', Polymer J, 1999 31 (12) 1231±1236. 17. Marshall JC, Cosgrove T, Jack K, Howe A, `Small-angle neutron scattering of gelatin/sodium dodecyl sulfate complexes at the polystyrene/water interface', Langmuir, 2002 18 (25) 9668±9675. 18. Su TJ, Lu JR, Cui ZF, Thomas RK, `Fouling of ceramic membranes by albumins under dynamic filtration conditions', J Membrane Sci, 2000 173 (2) 167±178. 19. Vermeer AWP, Norde W, `CD spectroscopy of proteins adsorbed at flat hydrophilic quartz and hydrophilic Teflon surfaces', J Colloid Interface Sci, 2000 225 (2) 394± 397. 20. Damodaran S, `In situ measurement of conformational changes in proteins at liquid interfaces by circular dichorism spectroscopy', Anal Bioanal Chem, 2003 376 (2) 182±188. 21. Kim J, Somorjai GA, `Molecular packing of lysozyme, fibrinogen, and bovine serum albumin on hydrophilic and hydrophobic surfaces by infrared visible sum frequency generation and fluorescence microscopy', J Am Chem Soc, 2003 125 (10) 3150± 3158. 22. Merrett K, Cornelius RM, McClung WG, Unsworth LD, Sheardown H, `Surface analysis methods for characterising polymeric biomaterials', J Biomater Sci Polym Edi, 2002 13 (60) 593±621. 23. Heavens OS, Optical Properties of Thin Solid Films, Dover, New York, 1965. 24. Born M, Wolf E, Principles of Optics, Pergamon, Oxford, 1970.

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25. Lu JR, Su TJ, Penfold J, `Adsorption of serum albumins at the air-water interface', Langmuir, 1999 15 6975±6983. 26. Lu JR, Su TJ, Thomas RK, `Structural conformation of bovine serum albumin layers at the air-water interface studied by neutron reflection', J. Colloid Interface Sci, 1999 213 426±437. 27. Lu JR, Thomas RK, `Problems in the analysis and interpretation of neutron reflection data', Nucl Inst Methods Phys Research, 1995 A354 149±163. 28. Lu JR, Perumal S, Powers E, Kelly J, Webster J, Penfold J, `Adsorption of b-hairpin peptides on the surface of water: a neutron reflection study', J Am Chem Soc, 2003 125 3751±3757. 29. Aggeli A, Bell M, Boden N, Keen JN, Knowles PF, McLeish TCB, Pitkealthy M, Radford SE, `Responsive gels formed by the spontaneous self-assembly of peptides into polymeric beta-sheet tapes', Nature, 1997 386 259±262. 30. Petka WA, Harden JL, McGrath KP, Wirtz D, Tirrell DA, `Reversible hydrogels from self-assembling artificial proteins', Science, 1998 281 389±392. 31. Hartgerink JD, Beniash E, Stupp SI, `Self-assembly and mineralization of peptideamphiphile nanofibers', Science, 2001 294 1684±1688. 32. Zhang SG, Holmes T, Lockshin C, Rich A, `Spontaneous assembly of a selfcomplementary oligopeptide to form macroscopic membrane', Proc Natl Acad Sci USA, 1993 90 (8) 3334±3338. 33. Vauthey S, Santoso S, Gong H, Watson N, Zhang S, `Molecular self-assembly of surfactant-like peptides to form nanotubes and nanovesicles', Proc Natl Acad Sci USA, 2002 99 5355±5360. 34. Lu JR, Perumal S, Hopkinson I, Webster JRP, Penfold J, Hwang W, Zhang SG, `Interfacial nano-structuring of designed peptides regulated by solution pH', J Am Chem Soc, 2004 126 8940±8947. 35. Xing Y, Zhang S, Olesen JT, Rich A, Guarente L, `Subunit interaction in the CCAAT-binding heteromeric complex is mediated by a very short alpha-helix in HAP2', Proc Natl Acad Sci USA, 1994 91 3009±3013. 36. Lu JR, Su TJ, Thomas RK, Penfold J, Webster J, `Structural conformation of lysozyme layers at the air-water interface studied by neutron reflection', J Chem Soc, Faraday Trans, 1998 94 3279±3287. 37. Lu JR, Su TJ, Howlin B, `The effect of solution pH on the structural conformation of lysozyme layers adsorbed on the surface of water', J Phys Chem B, 1999 103 5903± 5909. 38. Jackler G, Steitz R, Czeslik C, `Effect of temperature on the adsorption of lysozyme at the silica/water interface studied by optical and neutron reflectometry', Langmuir, 2002 18 (17) 6565±6570. 39. Su TJ, Lu JR, Thomas RK, Cui ZF, Penfold J, `The effect of pH on the adsorption of lysozyme at the hydrophilic silicon oxide-water interface, a neutron reflection study', Langmuir, 1998 14 438±445. 40. Su TJ, Lu JR, Thomas RK, Cui ZF, Penfold J, `The adsorption of lysozyme at the silicawater interface: a neutron reflection study', J Colloid Interface Sci, 1998 203 419±429. 41. Haynes CA, Norde W, `Structures and stabilities of adsorbed proteins', J Colloid Interface Surf, 1995 169 313±328. 42. Haynes CA, Sliwinski E, Norde W, `Structural and electrostatic properties of globular proteins at a polystyrene water interface,' J Colloid Interface Sci, 1994 164 394±409.

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43. Lu JR, Su TJ, Thomas RK, Rennie AR, Cubit R, `The denaturation of lysozyme layers adsorbed on the hydrophobed solid-water interface studied by neutron reflection', J Colloid Interface Sci, 1998 206 212±223. 44. Tanford C, Roxby R, `Interpretation of protein titration curves. Adsorption to lysozyme', Biochemistry, 1972 11 2192. 45. Lu JR, Thomas RK, Penfold J, Richards RW, `The segment density distribution profiles of polyethylene oxide adsorbed at the air-water interface, as studied by neutron reflection', Polymer, 1996 37 109±114. 46. Malmsten C, `Ellipsometry studies of protein layers adsorbed at hydrophobic surfaces', J Colloid Interface Sci, 1994 166 333±342. 47. Petrash S, Liebmann-Vinson A, Foster MD, Lander LM, Brittain WJ, Vogler EA, Majkrzak CF, `Neutron and X-ray reflectivity studies of human serum albumin adsorption onto functionalized surfaces of self-assembled monolayers', Biotechnol Prog, 1997 13 635±639. 48. Su TJ, Green RJ, Wang Y, Murphy EF, Lu JR, Ivkov R, Satija SK, `Adsorption of lysozyme onto the silicon oxide surface chemically grafted with a monolayer of pentadecyl-1-ol', Langmuir, 2000 16 4999±5007. 49. Prime KL, Whitesides GM, `Self-asssembled organic monolayer: model systems for studying adsorption of proteins at surfaces', Science, 1994 252 1164±1167. 50. Roth CM, Lenhoff AM, `Electrostatic and van der Waal contributions to protein adsorption ± comparison of theory and experiment', Langmuir, 1995 11(9) 3500± 3509. 51. Prime KL, Whitesides GM, `Adsorption of proteins onto surfaces containing endattached oligo(ethylene oxide): a model system using self-assembled monolayers', J Am Chem Soc, 1993 115 10714±10721. 52. Harder P, Grunze M, Dahint R, Whitesides GM, Laibinis PE, `Molecular conformation in oligo(ethylene glycol)-terminated self-assembled monolayers on gold and silver surfaces determines their ability to resist protein adsorption', J Phys Chem B, 1998 102 426±436. 53. Iwasaki Y, Fujike A, Kurita K, Ishihara K, Nakabayashi N, `Protein adsorption and platelet adhesion on polymer surfaces having phospholipid polar group connected with oxyethylene chain', J Biomater Sci Polymer Edn, 1996 8 (2) 91±102. 54. Murphy EF, Lu JR, Lewis AL, Brewer J, Russell J, Stratford P, `Characterisation of protein adsorption at the phosphorylcholine incorporated polymer-water interface', Macromolecules, 2000, 33 4545±4554. 55. Murphy EF, Lu JR, Brewer J, Russell J, Penfold J, `The reduced adsorption of proteins on the surface of phosphorylcholine incorporated polymer-water interfaces', Langmuir, 1999 15 1313±1332. 56. Murphy EF, Keddie JL, Lu JR, Brewer J, Russell J, `The adsorption of model proteins onto phosphorylcholine incorporated polymers studied by spectroscopic ellipsometry', Biomaterials, 1999 20 1501±1511. 57. Campbell EJ, O'Byrne V, Stratford P, Quirk I, Vick TA, Miles MC, Yianni YP, ASAIO J, 1994 40 M853. 58. Lewis AL, `Phosphorylcholine-based polymers and their use in the prevention of biofouling', Colloid Surf B, Biointerfaces, 2000 18 261±275. 59. Hayward JA, Durrani AA, Shelton C, Lee DC, Chapman D, `Biomembranes as models for polymer surfaces. 3. Characterization of a phosphorylcholine surface covalently bound to glass', Biomaterials, 1986 7 126.

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60. Armstrong J, Salacinski H, Mu Q, Seifalian AM, Peel L, Freeman N, Holt CM, Lu JR, `Interfacial adsorption of fibrinogen and its inhibition by RGD peptide: a combined physical study', J. Phys. Condens. Matter, 2004 16 S2483±S2491. 61. Su TJ, Lu JR, Thomas RK, Cui ZF, `Effect of pH on the adsorption of bovine serum albumin at the silica-water interface studied by neutron reflection', J Phys Chem B, 1999 103 3727±3736. 62. Su TJ, Lu JR, Thomas RK, Cui ZF, Penfold J, `The conformational structure of bovine serum albumin layers adsorbed at the silica-water interface', J Phys Chem B, 1998 102 8100±8108. 63. Lu JR, Murphy EF, Su TJ, Lewis AL, Stratford PW, Satija SK, `Reduced protein adsorption on the surface of a chemically grafted phospholipid monolayer', Langmuir, 2001 17 3382±3389. 64. Chapman RG, Ostuni E, Yan L, Whitesides GM, `Preparation of mixed selfassembled monolayers (SAMs) that resist adsorption of proteins using the reaction of amines with a SAM that presents interchain carboxylic anhydride groups', Langmuir, 2000 16 6927±6936. 65. Chapman RG, Ostuni E, Liang MN, Meluleni G, Kim E, Yan L, Pier G, Warren HS, Whitesides GM, `Polymeric thin films that resist the adsorption of proteins and the adhesion of bacteria', Langmuir, 2001 17 1225±1233. 66. McPherson T, Kidane A, Szleifer I, Park K, `Prevention of protein adsorption by tethered poly(ethylene oxide) layers: Experiments and single-chain mean-field analysis', Langmuir, 1998 14 176±186. 67. Herrwerth S, Eck W, Reinhardt S, Grunze M, `Factors that determine the protein resistance of oligoether self-assembled monolayers ± Internal hydrophilicity, terminal hydrophilicity, and lateral packing density', J Am Soc Chem, 2003 125 (31) 9359±9366.

13

Microgravimetry S V M I K H A L O V S K Y , University of Brighton, UK, V M G U N ' K O , Institute of Surface Chemistry, Ukraine, K D P A V E Y , University of Brighton, UK, P E T O M L I N S , National Physical Laboratory, UK and S L J A M E S , University of Brighton, UK

13.1 Introduction Microgravimetry is a generic name for sensitive analytical techniques based on measuring minor changes in the mass or density of an object. In the context of biomedical applications microgravimetry could allow the measuring of mass as small as that of an individual biomacromolecule. The most sensitive methods use the phenomenon of piezoelectricity, among which the most common technique is known as quartz crystal microgravimetry or microbalance technique usually abbreviated as QCM. Other microgravimetric methods, such as thermal gravimetry (TG), gas adsorption or dynamic contact angle analysis (DCA) are characterized by a lower sensitivity than QCM. The QCM method, which detects mass change at nanogram scale, is the most widely used technique for exploration of complex interfacial phenomena. It is also sometimes called quartz microBalance (abbreviated QMB), quartz crystal resonance sensor (QCRS) or quartz crystal immittance (QCI, QCM-
f-
R or QCM-
R). EQCM corresponds to electrochemical QCM whereas EQCN stands for electrochemical quartz crystal nanobalance. Combined measurements of frequency shifts
f and change in a dissipation factor (D) have been termed QCM-D. Deposition of a variety of surface functionalities on the QCM sensor surface provides specific, selective interaction with adsorbates. Other microgravimetric techniques cannot provide such selectivity and versatility of applications. Thus, the main focus of this chapter is on the use of QCM in chemical and biomedical analysis.

13.2 Quartz crystal microbalance technique More than 120 years ago Pierre and Jacques Curie discovered the electrical potential between mechanically deformed surfaces of certain crystals, such as Rochelle, or Seignette salt, quartz and tourmaline.1 This phenomenon is now known as the direct piezoelectric effect, as opposed to the converse piezoelectric

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effect, in which external electrical potential applied to a piezoelectric crystal causes its mechanical deformation.2 Piezoelectricity can only exist in a crystal that has no centre of inversion, as it requires a non-zero dipole moment to be formed in the crystal lattice upon mechanical deformation. For an unstrained crystal structure, the net dipole moment is zero, but if a mechanical force is applied across the crystal lattice, a configurational bond strain will occur and the net dipole moment will be nonzero. The dipole moment exhibits itself as an electrical field across the crystal structure. There are 21 of 32 crystal symmetry classes that do not have centre of symmetry, and 20 of them can exhibit the piezoelectric effect. For practical reasons however, only a few materials have been used in various applications. Quartz crystals dominate the market due to the combination of its chemical, mechanical and electrical properties described below. Quartz is a common naturally occurring crystalline material, however, natural quartz crystals of sufficient size and quality are relatively rare. Therefore, in most cases cultured quartz grown by hydrothermal synthesis is used.3 Natural (alpha) quartz belongs to the crystallographic class 32. It is an optically active mineral which exists in right- and left-handed forms. Most of the cultured quartz is the right-handed -quartz, although either form of quartz can be equally used to make resonators. The way in which quartz sheets (wafers) are cleaved from the main crystal determines their piezoelectric properties. The most commonly used AT-cut wafer is cut at 35ë150 to Z-axis. The advantage of the AT-cut resonator is that it has nearly zero frequency drift at temperatures around 25 ëC.4 Quartz has also low intrinsic losses, which means efficient conversion of mechanical into electrical energy or vice versa. Quartz has high a mechanical stress limit and a high insulation resistance, it is hard but not brittle, and it is chemically and thermally stable. When subjected to a mechanical deformation, a piezoelectric crystal behaves like a capacitor creating alternate charges on opposite faces; it can then generate an electric current. If a piezoelectric crystal is incorporated in an electric circuit, a very stable crystal controlled oscillator can be designed.2 Quartz crystals vibrate at a resonant frequency with minimal energy dissipation, which gained them wide application as nearly ideal oscillators in frequency controlling devices.5 Vibrational motion in the quartz crystal caused by the applied alternating electrical field generates acoustic standing waves. Their type and direction of propagation depends on crystal symmetry, angle of cut of the crystal and the arrangement of the electrodes used to apply the electrical field.6 Oscillating behaviour of a piezoelectric device can be explained and described using an electrical equivalent circuit model. The Butterworth-van Dyke (BVD) model proved to be useful in predicting frequency shifts in AT-cut QCM.7 Using this equivalent circuit, the properties of gases,8 liquids,9±13 interfaces,14 viscoelastic media15,16 and sensors17 have been investigated. Sinusoidal oscillations occur basically in one of two modes of the electromechanical resonance corresponding to the crystal equivalent circuit. These

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13.1 A simplified equivalent circuit model of a quartz crystal frequency resonator.2

modes are known as series and parallel resonant modes. In a simple mechanical/electrical description, the series-resonant mode depends on the `mass factor' L1 and the `spring constant factor' C1 of the crystal.18 At this frequency, inductor (inductance L1 ) and capacitor (capacitance C1 ) are seriesresonant and crystal impedance is essentially resistive (resistance R1 ). The parallel-resonant mode depends on the mass factor L1 and on the capacitance C0 of the electrodes having the crystal as a dielectric. At this frequency the terminal impedance of the crystal is resistive and very high (resistance  R1 ). To take account of resonances at overtones, other RLC series arms should be added in parallel to R1 , L1 and C1 (Fig. 13.1). C0 is real, or static capacitance; it represents the crystal in non-operational or static mode, plus capacitance of the electrode contacts and the crystal holder. L1 , C1 and R1 represent the crystal in an operational or motional state; inductance L1 represents inertia of the quartz plate, capacitance C1 represents its elasticity or stiffness and resistance R1 represents internal losses (dissipation of the oscillation energy) of the mechanical vibrating system.2,7 Modern QCM electrical circuits are much more sophisticated but they still use the Butterworth-van Dyke circuit as their background. If an alternating electric field and appropriate electronics are used, the crystal can be made to oscillate (Fig. 13.2) at its fundamental resonance frequency that depends on its geometry and thickness. A thicker AT-cut has a lower resonance frequency than a thinner crystal of the same shape, as in this case the thickness shear mode (TSM) or bulk acoustic wave (BAW) of vibration is used. In a TSM device the acoustic wave propagates through the crystal in a direction perpendicular to the crystal surface.19 Mass loading on the acoustic wave path alters the phase wave velocity causing a shift in the oscillation frequency of the crystal.20 AT-cut resonators operate in a frequency range of 1±30 MHz. To achieve a higher fundamental frequency an AT-cut wafer has to be so thin that it becomes difficult to manufacture. For example a fundamental frequency of 50 MHz requires a quartz resonator with a thickness of 33 mm.21 At such frequencies (30±250 MHz) overtones of the fundamental frequency are used. At frequencies below 1 MHz other cuts and modes of vibration are used. Depending on the mode of vibration used piezoelectric sensors are known as follows: TSM, surface acoustic wave (SAW), flexural plate wave (FPW), and shear horizontalacoustic plate mode (SH-APM) devices.6,21 A TSM device is undoubtedly

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13.2 Types of modes of crystal oscillations of acoustic sensors.3±6

popular as a sensor with a variety of biological applications despite a lower sensitivity than that of SAW device. SAW crystals oscillate at much higher frequency than TSM crystals (from 30 MHz to 250 MHz and even up to 1 GHz) and theoretically the sensitivity of the SAW technique could be higher by two orders of magnitude than that of TSM,22 but SAW devices require complex interface electronics and their sensitivity suffers upon immersion into a liquid.23 The influence of the liquid on QCM and SH-APM devices is less because the particle motion is parallel to the surface.6 It has been argued that the term QCM is not accurate and a more correct term `TSM resonator' or a BAW device should be used instead,22 but the first term is more commonly used in the literature and is used in this review. A quartz resonator has three main components, the quartz crystal substrate, a metal electrode on each face and two lead wires connecting it to the circuit. A quartz resonator used in QCM generally comprises a thin AT-cut quartz disc with a diameter between 3 and 25 mm, sandwiched between two metal electrodes which are used to establish an electric field across the crystal (Fig. 13.3). Crystals are usually produced in a circular form between 3 mm and

13.3 Quartz crystal with gold electrodes.

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25 mm in diameter depending on the application and required frequency. QCM electrodes can be made of any metal, but gold is the most common choice because of its inertness. Other metals such as silver or chrome are also used. A thin layer of metal is sputtered either directly onto the quartz surface or on a suitable underlayer of titanium, chromium or nickel.24 Typically, a 2 m layer of gold is formed on a 5±15 nm underlayer, the diameter and absolute thickness depending on the resonant frequency required. The electrodes trim the final crystal frequency by increasing the load on the surface. If the two electrodes are held at different potentials, an electric field results across the quartz crystal, i.e., in the `Y direction'. Because of the piezoelectric properties of quartz, such an electric field in the `Y direction' couples to shear motion `around' the Z-axis, and vice versa. The end result is that shear waves in the quartz, in which the mechanical displacement occurs in the `X' direction, also called the electric axis, are coupled to the voltage between the electrodes. In 1959 Sauerbrey25 showed that the frequency shift of a quartz crystal resonator is directly proportional to added mass, which was a major step towards developing a new quantitative tool to measure very small masses, i.e., the quartz crystal microbalance as an ultra-sensitive mass sensor. The measured frequency is dependent upon the combined thickness of the quartz wafer, metal electrodes and material deposited on the QCM surface. The larger the mass added, the greater is the frequency decrease. Because the resonance is very sharp, high precision frequency measurements allow the detection of minute amounts of deposited material, as small as 100 picograms on a square centimetre. Mass changes occurring at the QCM surface result in frequency changes according to the Sauerbrey equation 2f 2
m
f ˆ ÿ p0  ˆ ÿ2:26  106 f02
m=S S q q

13:1

where f0 is the resonant frequency of the unloaded quartz crystal,
m is the adsorbed mass, q is the specific density of quartz (2.648 g cmÿ3), q is the shear modulus of quartz (2.947  1011 g cmÿ1 sÿ2), and S is the surface area of the sensor in cm2. In high sensitivity QCM devices using an overtone (higher than the fundamental) frequency, eqn 13.2 should be used:
f ˆ ÿ

2fn2
m p nS q q

13:2

where fn is the frequency corresponding to an odd harmonic (n ˆ 3, 5, 7 . . .). p Factor Cf ˆ ÿ2f02 =…S q q † is known as the quality factor; it is a property of the quartz crystal independent of the properties of the film, which is a very strong advantage of QCM mass sensors. Assuming that the film is uniform, its

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thickness df is calculated by dividing the film mass
m determined from eqn 13.1 by the film density f . df ˆ


m f

13:3

For a 5 MHz AT-cut quartz crystal it is 56.6 Hz gÿ1 cm2 at room temperature.26 At a larger film thickness its properties such as density and elasticity have to be taken into account.27

13.2.1 QCM measurements in liquids The Sauerbrey equation is only strictly applicable to uniform, rigid, thin-film deposits, and if the crystal oscillates in the first harmonic. It was derived from the assumption that the added film is so rigid and thin that it can be treated as an extension of the thickness of the quartz crystal with the same properties. In many cases the equation is not valid. For example, when the added mass (i) is not deposited rigidly on the electrode surface, (ii) slips on the surface or (iii) is not deposited evenly. The Sauerbrey equation does not provide a satisfactory model for frequency changes in liquids because of the change in viscosity, density and concentration of solutes. Crystal oscillations bring the adjacent liquid layer into motion and the latter interferes with the former making the relationship between
f and
m non-linear and difficult to interpret. A rougher crystal surface may further increase
f caused by the liquid.28 Due to this, the QCM was for many years just used as a mass detector for thin film deposition from gas phase. Not until the beginning of the 1980s was it realized that a quartz crystal can be excited to a stable oscillation when it was completely immersed in a liquid.29,30 Pioneering work in quantitative liquid phase QCM measurements was done by Kanazawa and Gordon,31 who showed that the frequency shift of a QCM in a liquid is proportional to the square root of its density and viscosity as follows: r `  ` 3=2
f ˆ ÿf0 13:4 q q where
f is the measured frequency shift, ` and ` are the density and viscosity of liquid in contact with the crystal, respectively. Despite much higher viscous loading in liquids, the QCM response is still extremely sensitive to mass changes at the solid/liquid interfaces. QCM sensors are used in direct contact with liquids and/or visco-elastic films (Fig. 13.4) to assess changes in mass and visco-elastic properties of the studied systems.31,32 Among the advantages of quartz crystal sensor-based methods are their robustness and ability to provide non-invasive, label-free and sensitive measurements.33 Although the original theory of QCM techniques has suggested energy losses into the surrounding media and would set the limit of QCM analysis to films

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13.4 Resonator geometry.2

with thickness below 1 m, it has been shown that QCM methods work with surface coatings of a thickness in the range of 2±3 m34 and recently new circuitry based upon auto-gain principles has been developed, thought to be capable of further extending this working range.35 The Sauerbrey equation overestimates the mass of protein films adsorbed from aqueous solutions because the hydrated proteins add water molecules to the measured mass and the aqueous solution further dampens the frequency.31 In the case of adsorption of a film with mass ma on the crystal from a liquid, oscillation frequency shift can be described by eqn 13.5:   ` ` 0:5
f ˆ ÿf01:5 ÿcma …1 ÿ x†; 13:5 q q where c is a constant, and x is the viscous correction factor dependent on mechanical properties of an adsorbed film and an aqueous solution.36 If this film (with the thickness df and the specific density f ) is relatively rigid and the liquid is Newtonian, the load from this layer and the liquid load are approximately additive:37   ` ` 0:5 2f02 f hf
f ˆ ÿf01:5 ÿ 13:6 q q …q q †0:5 A solution in contact with a QCM gradually changes its properties such as density and viscosity upon film formation; these changes also have an effect on the measured frequency31,38,39 Hence, measuring frequency shift alone is not sufficient for gaining correct information about properties of the adsorbed layer. Another parameter, which corresponds to resistance R1 in the BVD equivalent circuit, is known as damping, or quality factor (Q), or dissipation factor D (D ˆ 1=Q). It measures the efficiency of conversion of mechanical energy into electrical energy in QCM. Experimentally it is measured by recording QCM

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voltage decay when the voltage is switched off.8 The decay curve is described by an exponentially damped sinusoid:40 A…t† ˆ A0 eÿt= sin …!t ‡ '†

13:7

where A is the amplitude, t is time,  is the constant and ' is the phase. By numerical fitting of the voltage decay curve both
f and D are determined. Measurements of frequency change
f …t† and dissipation energy D(f,t) are the two most common parameters associated with the interaction of adsorbed compounds, the surroundings and the sensor surface. Interpretation of QCM-D data relies on the model used for curve fitting.41 It can be argued, therefore, that because D is a function of f, other parameters such as resistance and the maximum oscillation amplitude should be measured to improve the interpretation of QCM data.36 For the Butterworth-van Dyke circuit the series resonance frequency is given by eqn 13.8:22   1 C0 D2 13:8 f ˆ p 1 ‡ 2C1 2 L1 C1 where the dissipation factor D ˆ R1 =!L1 . It has been shown experimentally that the maximal oscillation amplitude A0 is linearly proportional to Q for quartz crystals immersed in liquid, provided no film deposition occurs.39 This is not the case when adsorption on a quartz crystal surface takes place and the theoretical modeling becomes complicated. A summary of the parameters measured by QCM techniques is given in Table 13.1.39 Another electrical analogue description of a layered compound resonator's behaviour models the system as a transmission line.42 Many authors have chosen this approach to relate observed impedance/admittance spectra to overlayer and interface properties.43±47 Although electrical studies could make valuable contributions to the investigation of film properties, they can often mask physical insight into the detailed nature of loading mechanisms.48,49 Models which take into account the correlation between the motion of the quartz and the overlayer, as well as the influence of the electronic circuit on the electrical state of the whole system are of fundamental importance in interpreting the results obtained with the help of QCM.18 Table 13.1 Comparison of the characteristics of quartz and GaPO4 (refs 52, 53) Parameter Phase transition (ëC) Piezoelectric sensitivity (pC/N) Temperature compensate cuts Coupling coefficient (%) Frequency constant (MHz mm)

Quartz

GaPO4

573 2.3 Yes 8.8 (AT) 1.66

970 4.5 Yes 16 (Y-16.5ë) 1.27

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13.2.2 Electrochemical QCM The use of QCM in a liquid medium together with electrochemical methodologies has led to the development of a new technique ± EQCM.50,51 The quartz crystal sensor face in contact with the electrolyte constitutes the working electrode in a three-electrode electrochemical cell also comprising a reference electrode and a counter electrode. In solid state voltammetry, mass changes occur at the electrode surface as oxidation and reduction occur. The EQCM setup allows simultaneous measurement of mass and current in a voltammetric experiment. The first in situ application of QCM to electrochemistry was made by Nomura and co-workers to study Cu(II) and Ag(I) electrodeposition. 29,52 Most commonly, in EQCM analysis electrochemical mass deposition measured by QCM is related to the charge transfer measured by an electrochemical method. Quantitative analysis is based on using both the Sauerbrey eqn 13.1 and Faraday's Law, which link
f to the transferred charge q. It requires carrying out experiments in conditions when the eqn 13.1 is valid. A combination of eqn 13.1 and Faraday's Law gives a linear relationship between
f (in Hz) and q (in Coulombs) as in eqn 13.9: 106
M
Cf
q 13:9 nF
S where M is the molar mass of the depositing species (in grams), Cf is the quality factor in eqn 13.1, n is the number of electrons transferred, F is Faraday's constant and S is the effective surface area of the QCM electrode in cm2. Cf and S are determined by calibrating the EQCM using some well characterised electrochemical reaction such as silver electrodeposition on gold electrode.53 Most QCMs are driven by known oscillator circuits,40 which only allow monitoring of resonant frequency changes for low damping media. In a liquid medium, mass change measurements must be correctly interpreted by taking into account the viscous/visco-elastic properties of the medium.54 A nonconventional driver has been proposed for EQCM applications when damping needs to be taken into consideration.16 It consists of a controlled gain voltage amplifier with positive feedback, the voltage gain being determined by a potentiometer. By increasing the potentiometer resistance, P, just to the start of oscillations, the amplifier oscillates in a series-resonant mode with P ˆ R1 , thus allowing the monitoring of both resonant frequency changes
f and R1 values. Consequently, the power of the modified EQCM16 resides in its ability to make measurements of visco-elastic properties of films (for example, polymer films and sol-gel phase transformations47,54). One of the first applications of this setup demonstrated that
f depends on
R1 , this dependence being related to C0 and to the phase angle added by the driver. Measurements of
f , with corrections for
R1 were published by FruboÈse et al.;54 an application of the modified EQCM for the electrodeposition
f ˆ

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of polypyrrole films was reported.55 It should be stressed that the same dependence should be taken into account when conventional oscillator circuits are used. The main difficulty is that they do not give the value of R1 , and the corrections for
f measurements become impossible. Conventional use of EQCM is to measure double layer structure changes at gold electrodes when no visco-elastic effects are detected. Species in the outer and inner Helmholtz layer, ion solvation, ion pair formation and adsorbed ions also contribute to the measured frequency shifts.56 EQCM detectors can achieve high frequency resolution (0.1 Hz) and mass resolution (0.05 ng).22 A new flow-through cell was designed for simultaneous electrochemical and QCM analysis.57 It comprises three electrodes ± reference, auxiliary and a working Au/quartz electrode, the latter being used also for piezoelectric mass detection. This cell worked in conjuction with flow analytical techniques such as flow injection analysis (FIA) and high performance liquid chromatography (HPLC). To determine the flow pattern, the cell was tested by analysis of 1 mM K4Fe(CN)6 in 0.5 M KNO3 using 5 or 10 MHz resonant frequency quartz crystals. The system operated in a nearly radial-flow thin-layer mode and was successfully tested in analysis of water-glycerol mixtures, electrodeposition of Ag from 1 mM AgNO3 in HClO4, detection of pesticides using a molecularly imprinted polymer film and nucleotides using self-assembled monolayer films of DNA.

13.2.3 Additional parameters Visco-elastic materials such as polymer (protein) films and a viscous liquid medium influence the frequency shift of the quartz crystal, and it is difficult to distinguish their contribution to
f from the mass contribution. Analysis of the frequency shift and the resistance change of QCM coated with a visco-elastic film and operating in a viscous liquid provides additional information for advanced signal interpretation. Lucklum and Hauptmann suggested a method based on the `Acoustic Load Concept' (ALC) for sufficiently accurate estimation of the complex shear modulus of a coating.58 With this value, impedance approximation can be applied to calculate film thickness more accurately than with those equations based on an assumption of thin, rigid films. The acoustic load impedance is calculated from electrical impedance. In the ALC generalized (complex) values of the shear modulus and the acoustic impedance and (imaginary) interfacial layers are used. The latter may contain only characteristic impedance and a phase shift of the acoustic wave. This concept allows one to calculate effective values directly from experimental results. It furthermore allows the computation of these effective values from different physical models, which do not need to be one-dimensional.59 Impedance and shear wave resonance were used to study ligand-receptor interactions at functionalized surfaces and at cell monolayers.60 The roughness

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of the crystal surfaces can affect the liquid layer in contact. An impedance analysis study carried out by Martin on crystals of different levels of roughness showed that the frequency shift increases with growing roughness.61 Provided the roughness was less than 0.15,  being the decay length of the shear wave generated by the oscillating surface, the approximation of the surface as hydrodynamically smooth remains valid. By replacing conventional gold electrodes with porous gold electrodes, the effective adsorbing surface area and hence sensitivity was enhanced by a factor of 40 with no sacrifice of the mechanical quality factor or the stability in the resonant frequency.62 Increase of the ionic strength leads to a significant frequency drop and may even go below design limits, stopping crystal operation.13,63 This effect is caused by an electrical short-cut created by the electrolyte solution acting as a conductor between the electrodes on opposite faces of the QCM crystal. This problem is, however, easily overcome by immersing only one face of the quartz crystal in the liquid phase.

13.2.4 Sensors with other piezoelectric crystals Non-quartz sensor crystals, such as GaPO4 possessing high sensitivity and thermal stability, can be used in media or temperature conditions inappropriate for quartz crystals. It has a crystal structure homeotypic to quartz. Gallium orthophosphate has higher piezoelectric sensitivity and thermal stability than quartz, and its crystals can be cut so that the temperature compensation would be achieved for any temperature range within its thermal stability.64 The resonance frequency shows very good reproducibility during multiple heat-up cycles, for example it shifts only by 30 ppm in the range from 350 ëC to 650 ëC. GaPO4 plano-convex resonators with a diameter of 10 mm were tested as piezoelectric microbalance sensors. They showed high Q-values up to 1.2 million at fundamental frequency of 6 MHz in vacuum. GaPO4 can be used with standard quartz holders and Au electrodes. Among other piezoelectric materials that have been tested as a replacement for quartz, are a few ceramic piezoelectrics such as PZT (lead-zirconate-titanate, solid solution of PbTiO3 and PbZrO3), lithium niobate, lithium tantalate and also ZnO and PVDF (polyvinylidene-fluoride).65±69 They are seldom used as mass sensors and their advantages over quartz have still to be proved.

13.3 Analytical applications of QCM 13.3.1 QCM as an analytical technique Sensitivity of QCM in measuring mass changes is at least 100 times higher than that of an electronic balance with a sensitivity of 0.1 mg.70 This means that QCM is capable of registering objects as small as a fraction of a monolayer or a

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single layer of atoms. High sensitivity and instrumental possibility of the realtime monitoring of mass changes make QCM a very attractive technique for a large range of applications. The development of QCM systems for use in fluids or with visco-elastic deposits has dramatically increased interest in this technique. The major advantage of the QCM application in liquid systems is that it allows a non-invasive label-free detection of molecules with a sensitivity matching radiolabelling and fluorescent-labelling techniques. Thus, in the last decade, QCM as a new analytical method for in situ investigation of interfacial processes including electrode processes (EQCM) has emerged.50,57,71±74 A partial and already long list of QCM applications as an analytical technique (in no particular order) is as follows: (i) thin film thickness monitoring in thermal, e-beam, sputtering, magnetron, ion and laser deposition; (ii) electrochemistry of interfacial processes at electrode surfaces; (iii) interactions of DNA and RNA with complementary strands; (iv) specific recognition of protein ligands by immobilized receptors and immunological reactions; (v) detection of virus capsids, bacteria, mammalian cells; (vi) adhesion of cells, liposomes and proteins; (vii) biocompatibility of surfaces; (viii) biofilm formation; (ix) creation of selective surfaces; (x) lipid membranes; (xi) polymer coatings; (xii) reactive surfaces; (xiii) gas sensors; (xiv) immunosensors; (xv) thin film formation; (xvi) Langmuir and Langmuir-Blodgett (LB) films; (xvii) self-assembled monolayers (SAM); (xviii) polyelectrolyte adsorption; (xix) spin coating; (xx) bilayer formation; (xxi) adsorbed monolayers; (xxii) surfactant interactions with surfaces; (xxiii) dissolution of polymer coatings; (xxiv) molecular interaction of drugs; (xxv) cell response to pharmacological substances; (xxvi) drug delivery; (xxvii) relative humidity, and (xxviii) in situ monitoring of lubricant and petroleum properties.

13.3.2 Non-selective mass measurement The quartz crystal microbalance technique is essentially a non-selective tool, as it responds to a mass load change, and for many applications it requires additional surface modification to make it more selective (see section 13.3.3), however even a non-selective QCM has been used in studying adsorption, metal deposition and film growth.75 Uncoated QCMs as chemical sensors Early applications of the QCM technique involved well-documented measurement of metal deposition in high-vacuum metal evaporators, and it is still widely used to determine the layer thickness.76 It allows for real-time, rapid measurement of film thicknesses with Angstrom resolution. QCM is also suitable for detection of volatile substances in vacuum, air or other gases.77 QCM was used as a mass detector of nitroaromatic compounds separated in a

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chromatographic column.78 The sensitivity of GC-QCM technique was similar to that of GC-MS, and it was proposed as an inexpensive alternative to the latter. Original use of QCM as a humidity sensor is described.79 An AT-cut quartz plate vibrating at 5 MHz was mounted directly on the Peltier element. The quartz plate was cooled down to the dew point by the Peltier element, and the mass of the water condensed from air on the QCM was measured. The sensor was able to measure relative humidity between 20% and 95%. Advances in QCM methodology in the last decade now allow for dynamic measurement of minute mass changes at surfaces, thin films and electrode interfaces prepared on a quartz crystal, while the surface is immersed in a liquid. Combining QCM technique with electrochemical instrumentation allows simultaneous measurement of mass and electrochemical variables such as electrochemical potential, current and charge. A QCM sensor based on 30 MHz AT-cut shear mode crystal was tested in gas and liquid phases.12 As expected, higher fundamental frequency resulted in higher sensitivity of the device. It was used for EQCM analysis of copper electrodeposition at the quartz crystal surface. Although thinner and hence more fragile quartz crystals are required, the benefits of using higher operating frequency include better performance characteristics than 5 or 10 MHz crystals typically used in liquid media applications. Investigation of polynaphthylamine film growth on the Pt electrode of a QCM sensor in acetonitrile has been studied.80 Combining QCM with fibre optic reflection spectroelectrochemistry, simultaneous measurements were performed. By immersing one side of the quartz crystal in an aqueous solution whilst keeping the other in the air, the behaviour of Langmuir films formed at air/water interface was studied.81 Attachment and detachment of LB films has been measured using quartz crystal surface as a substrate.82 Adsorption of phospholipids from a concentrated liposome solution and stabilized oil-in-water emulsion was studied using QCM-D at 5 MHz.83 The third overtone was also analyzed as being more sensitive to the close-to-surface layer than the fundamental frequency. Use of QCM-D and analysis of data obtained with two frequencies produced a very detailed picture of the adsorption behaviour of the systems studied. It was found that liposomes rapidly adsorbed on a gold surface, and created a phospholipid bilayer, to which liposomes slowly continued to attach. This attachment was rather loose as rinsing the substrate with water led to liposome desorption, whereas the phospholipids film remained on the surface. A rapid initial step of emulsion adsorption was followed by slower surface spreading of the droplets. Although the droplets were replaced from the surface by water upon rinsing, a monolayer was retained by the surface. Gold has high affinity with sulphur-containing compounds; this property of the gold electrode surface in QCM was used to study self-assembly of a synthetic clay containing thiol groups from chloroform solution.84 Thiol adsorption in a flow injection system was studied using QCM with impedance

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analysis. It was found that the presence of electrolytes interferred with mass detection, and experimental error could be as high as 70%.85 QCM with gold electrodes was used to study adsorption/desorption of a number of short chain alkylthiols and alkyl sulfides from gas phase.86 The results showed that the instrument was capable of distinguishing between different adsorbates as they had different affinity with gold. An interesting example of a QCM application was to study evaporation of volatile organic substances.87 Evaporation of a sessile droplet of light alcohol from the oscillating quartz crystal surface was monitored by measuring a decrease of the mass load and hence an increase of the oscillating frequency. Another unusual application and a place as well, of the QCM technique was with the Pathfinder mission, which used piezoelectric mass change detection to measure how fast dust settles on the surface of Mars.88 This was not the first time that NASA utilized QCM, which was also put on board Discoverer 26 to measure the surface erosion for the Atlas Missile Program in 1961. The aggregation of starburst poly(amidoamine) dendrimers bearing either 64 carboxylic or amine peripheral groups from aqueous solution on the surface of bare and thiol modified gold electrodes was studied using the QCM technique.89 Thiols also contained either carboxylic or amine groups that were exposed to the solution. The electrostatic interaction between the dendrimer molecules and the modified gold surface promoted much larger surface coverage than on bare gold. Coated QCM as chemical sensors The example given above explains the reason why a coated rather than an uncoated quartz crystal sensor is used in the majority of applications. Coating significantly increases the selectivity and sensitivity of QCM analysis.90 Coating of the crystal surface may serve two purposes: (i) it can be an object of analysis itself, for example, to study an interaction between the material of coating and other substances, or (ii) it can be used to specifically interact with the analyte in the liquid or gaseous medium. In the latter case the nature and choice of the coating is very often the same as the nature and choice of the stationary phase in gas chromatography.91±94 As gas sensors of volatile organic compounds, QCM detectors are used for designing `electronic noses', or e-noses.27,95±98 Although it should be said that creating a sensor that matches human nose sensitivity still remains the subject of laboratory research, coated QCM-based sensors for monitoring air pollution have been commercialized.22,99 They are used to detect volatile organic compounds, environmental pollutants and as chromatography detectors.30 Mostly these coatings are made from polymer films or stationary phases used in gas liquid chromatography.93,94 A very high frequency (VHF)-band QCM gas sensor operating at 77 MHz was reported.100 It has become possible to make very thin quartz plates using anisotropic etching, so that they are supported by thicker quartz plates around

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Surfaces and interfaces for biomaterials

electrodes. The sensor and the oscillator circuit were placed on the opposite sides of the printed circuit board to make the wiring length as short as possible because of the high resonant frequency. The VHF sensor was coated with a free fatty acid polyester used as a stationary phase in gas chromatography. For the detection of acidic gases such as CO2, NO2 and SO2, polymers and low molecular substances with basic properties such as N,N,N',N'-tetrakis-2hydroxylethyl ethylenediamine101 and aminated polystyrene102 have also been used as coatings. Respectively, polymers containing carboxylic groups were used to increase the selectivity of the QCM method towards detection of organic amines in water.103 It was found that reversible changes in resonance frequency of the coated QCMs were characteristic of acid-base interaction and hydrogen bonding, whereas quasi-irreversible frequency changes were found for saltcomplex forming reactions between the solute and the polymer film. The relative selectivity of the analytes decreased with the increase of the hydrophobic part of the solute molecules in aqueous phases when acid-base interaction between the coating materials and analytes were employed. Polyvinylpyrrolidone coated QCM also showed selectivity towards adsorption of ammonia and short chain primary aliphatic amines from air.104 Its shelf life was at least five months without loss in sensitivity and performance. Often one coating cannot provide sufficient selectivity for all analytes, especially in such complex multi-component systems as perfumes and flavours. In this case an array of QCMs, each topped with a different thin film which absorbs a particular set of chemicals is used.93,105±108 When these chemicals are present in the environment, they are absorbed increasing the mass of the QCM and decreasing its resonant frequency. The pattern of sensor frequency shift gives information about the composition of the analyzed odours. Polymer films have been coated on the QCM resonator surface by electrodeposition,109 self-assembly,110 spraying,111 casting from solution,104 spin coating,112 glow discharge plasma polymerization,113 photopolymerization114 and molecular imprinting.115 The Langmuir-Blodgett coating has certain advantages over casting methods, because the thickness of LB films, as well as their composition and molecular orientation can be better controlled. LB film coating is obtained by transferring monolayers from an air-liquid surface to the QCM surface.116,117 QCM with immobilised LB films of poly(3-butoxythiophene) (P3BOT) mixed with stearic acid (SA) detected vapours of low molecular alcohols and chlorine-substituted methanes; the sensitivity of the sensor increased with the number of LB layers.117 Polymerization in the presence of analyte molecules ± molecular imprinting has been used to design QCM for analysis of drugs and other small organic molecules in gas phase as well as in a liquid: acetaldehyde,118 caffeine,119 volatile organic compounds,115 nandrolone,112 aged rape seed oil,120 sorbitol121 and glucose.122 QCM sensors coated with molecularly imprinted polymers

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(MIP) have greater sensitivity and higher selectivity than sensors coated with non-imprinted polymers. MIP-coated QCM have been reported to have chiral selectivity in detecting S-propranolol,123 L-menthol124 and L-serine125 with good detection limit in the ppb range. Among other classes of coatings, zeolites were used for humidity, NO, SO2 and volatile organic substances sensing.126±128 Zeolite structure and pore size influence the sensitivity of the QCM sensor.127 A QCM sensor for dissolved oxygen was designed by self-assemblying layers of tris(2,2'-bipyridyl dichlororuthenium).129 A powder of another transition metal complex ± [Pt(CN-cyclododecyl)4][Pt(CN)4] was incorporated in a hydrophobic vapour permeable polymer and the composite film attached to the QCRS surface. The platinum complex can reversibly bind water molecules, and this reaction was used to measure humidity.130 Attempts to use carbon coatings have been reported. They include carbon nanotubes,131 fullerene C60 and its derivatives,132 graphite133 and carbon black±polymer composite134 for detection of alcohol and other polar organic substances in gas phase at room temperature. Boron-doped diamond films proved to be suitable for fabricating a QCM electrode and using it in EQCM by examining hydrogen and oxygen evolution.135 To detect heavy metal ions in solution, QCM sensors were modified with polymers containing functional groups that strongly bind these ions via formation of surface complexes, such as polythiophenes (Hg2+ with sulphur)136 or poly(acryloyl morpholine) (Cu2+ with nitrogen).137 QCM coated with a crownether containing co-polymer of acrylic acid with ethylene glycol dimethacrylate demonstrated selective response to K+ ions with a detection limit of 0.4 ppm in the presence of Li+ and Na+.138 Binding between Group I and Group II metal ions and oxacalix[3]arenes was studied using QCM.139 The macrocyclic ligands were adsorbed on the gold-plated surface of a quartz crystal sensor modified with a selfassembled monolayer of 11-mercaptoundecanoic acid. Frequency shifts detected by QCM showed a strong preference for binding Na+ compared to K+ or Ca2+, which correlated with the geometry of the macrocycle's central cavity. QCM coated with p-tert-butylcalix[4]arenetetrathiolate monolayer showed selectivity towards alkylbenzenes,140 whereas serine-containing dicyclodipeptide-bearing calix[4]arene coatings were capable of enantiomeric recognition of (R)-methyl lactate rather than its (S)-enantiomer,141 suggesting that these ligands may be used to construct highly selective quartz piezoelectric sensors by adjusting their cavity to the analyte of interest. Ion-selective QCM sensors were also made by forming crystals of scarcely soluble inorganic salts such as CdS and CuS, on the quartz crystal surface. These salts have properties of ion-selective solid membranes due to preferential adsorption of certain ions.142 Similarly, selective orthophosphate QCM sensors were made by immobilizing insoluble phosphates of trivalent Ce, Cr and Bi on quartz crystal surface.143 Microcrystals of phosphates were embedded in a layer of the silicone rubber. At pH 7.0 orthophosphate ions were detected in the range from 10ÿ6 to 10ÿ2 M.

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QCM devices coated with monomeric and polymeric phthalocyanine films and crown ethers have been tested as sensors for the detection of NO2 and a number of organic substances such as alcohols, terpenes and aromatic molecules.144,145 As expected, the results vary significantly depending on the type of film used because the properties of phthalocyanines depend on the type of the central metal ion, attachment and position of the side groups and substituents. Monitoring of hydroxyapatite (HA) interaction with a tea staining solution using QCM has been reported.146 To mimic tooth enamel, a crystalline HA layer was electrodeposited on the surface of the quartz crystal sensor precoated with a phosphate-terminated polymer. HA-coated QCM detected adsorption of staining substances from an extract of tealeaves passed over the sensor. This system was used to model dental stain formation and to study efficiency of stain removal by tooth-whitening products. One of the main problems with using coated QCM is the stability of the coating. To ensure stronger adhesion between the film and the gold electrode, sulphur-containing polymers have been used.147,148 1-Octadecanethiol was selfassembled on a gold-coated AT-cut quartz, thus creating a hydrophobic surface used to study the adsorption of lipid vesicles.149 A smectite-type clay with thiol groups was self-assembled on the gold surface of the QCM electrode from the chloroform suspension.150 Stability and reactivity of six amine functionalized coatings made from 4aminopyridine, 4-aminothiophenol, 2-aminoethanethiol, 3-amino-5-mercapto1,2,3-triazole, -aminopropyl-triethoxysilane and polyethyleneimine was studied via reaction with succinic anhydride.151 Only the polymeric film showed both stability and sensitivity probably because other substances did not form strong bonds with the surface. Comparison of the stability of different coatings such as phospholipids and several stationary phases used in gas chromatography showed that lipid films were stable whereas low molecular GC stationary phases PEG 200, octadecane and DBP were not; the best stability was achieved when a lipid film was blended with polyvinylchloride and its plasticizer DOPP.152 Another problem is the applicability of the Sauerbrey equation and in general of the rigid film model. It has been shown that films of Prussian blue with thickness up to 0.15 m can be studied by measuring pure mass change, whereas ignoring the visco-elastic behaviour of thicker films leads to substantial mass errors.153

13.3.3 Bioanalytical applications of QCM It is not always possible to distinguish between chemical, biological and immunosensors because the difference between them is rather conditional; here we follow a convention that a biosensor has a biomolecule as an active component of the sensing device whereas an immunosensor involves an antibody ± antigen interaction to measure the analyte.

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QCM biosensors A biosensor is a device with an active sensor material of biological origin in close proximity to a transducer prepared, e.g., by immobilizing biomolecules onto the surface of the transducer.17,154±157 There are four widely used techniques for the immobilization of a species onto an electrode: (i) adsorption; (ii) chemical bonding directly to the electrode or covalent coupling to an appropriate inert support attached to the electrode, (iii) cross-linking of biomacromolecules to the electrode, and (iv) entrapment at the electrode surface using an appropriate polymer. Grafting of small functionalities, such as nucleotide bases, aminoacids and calix[n]arenes, onto the sensor surface provides specific interaction of the sensor with target compounds.158 Sensor surface coverage with functional polymers, lipids, proteins, DNA and enzymes allows for direct monitoring of biospecific interactions. Real-time monitoring of interactions between bioactive species such as DNA, proteins, lipids, drugs and metal ions in natural environment is of interest for bioscience, and interaction of biomacromolecules with artificial surfaces (biomaterials) is of essential interest for many medical and biochemical applications.32,159 A variety of methods: spectroscopic (fluorescence, FTIR, circular dichroism and XPS), optical (ellipsometry and reflectometry), differential scanning calorimetry, adsorption, radiolabelling and sensor techniques are used for these purposes.160 Immobilization of certain receptors onto the electrode deposited at a quartz crystal, in active orientation, followed by exposure to studied ligands, with frequency monitoring, allows investigation of any interaction between receptor and ligand as a change in the resonant frequency of the crystal in real time, and so the kinetics of the binding processes can be measured. Guilbault154 developed a formaldehyde detector using an enzyme coated quartz crystal. Muramatsu et al.161 used crystals covered by lipid layers containing asolectin and cholesterol for the recognition of various analytes. A QCM biosensor combined with FIA system was used for studying interactions between a protein and small molecules.162 Two sulfa-drugs, sulfamethazine (SMZ) and sulfamethoxazole (SMO) were immobilized on Au electrodes of the quartz sensor via self-assembled layer of dithiothreitol. The binding of various proteins such as human, goat and mouse IgG, trypsin and chymotrypsin to the immobilized drugs was studied. The QCM analytical technique was capable of distinguishing specific protein ± drug interactions despite similarity of the drug structures, and the quantitative parameters of these interactions were calculated. A quartz crystal resonance sensor with immobilized human serum albumin (HSA) was used for studying interaction of warfarin and diazepam with HSA.35 This warfarin-HSA binding had been previously investigated by a number of techniques, and protein conformational shifts were expected to occur as a result

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of binding of the relatively small molecule to the macromolecule. The results showed that the sensor response in frequency and resistance depended on the particular method used to attach the protein to the device surface. The latter involved gold-sulphur surface chemistry through interfacial linking molecules such as thioctic (-lipoic) acid, cysteamine, 11-mercaptoundecanoic acid or dithio-bis-(succinimidylpropionate) QCM sensor with immobilized bovine serum albumin did not respond to either drug. Molecular imprinting of polymers with large biological objects such as enzymes, viruses and cells was used to design QCM biosensors.120 Cell imprints were made in polyurethanes by pressing `cell stamps' on a pre-polymerized polyurethane layer, the material was cured overnight and the cells were removed with hot water. Viruses were stamped onto pre-polymerized mixtures of methacrylic acid (MA), styrene and divinylbenzene (DVB) with an initiator and then cured under UV light. The viruses were removed with a hot aqueous detergent solution. Imprinting with proteins required use of an oligomeric mixture of MA, styrene and DVB cured afterwards under UV irradiation. Subsequently MIP-coated QCM sensors were used to detect yeast cells, tobacco mosaic virus and trypsin. The MIP-sensors showed much higher selectivity of detection than QCM coated with non-imprinted polymer films.120 Identification of DNA sequences in clinical, forensic, food and environmental samples has become a powerful tool in detection and diagnostics of infectious and hereditary diseases, biological species, and bacteria and viruses.77,163 Using the high affinity of the biotin ± streptavidin interaction, a layer of streptavidin was covalently attached to a thiol/dextran modified Au electrode surface, to which biotinylated oligonucleotides were immobilized on a streptavidin coated QCR (quartz crystal resonator) sensor. The DNA QCM sensor has been coupled to the polymerase chain reaction (PCR) and applied to detecting a bacterial toxicity factor aerolysin, and polymorphism of human apolipoprotein E through the analysis of amplified DNA samples extracted from different samples and amplified by PCR. A DNA piezoelectric sensor has been also used for the detection of genetically modified organisms (GMOs).164 The selectivity of analysis was achieved by immobilizing particular types of oligonucleotides, complementary to foreign DNA sequences present in genetically modified food. Oligonucleotide immobilization was realized by two different techniques: (i) either via interaction of the biotinylated ligand with the streptavidin-coated piezoelectric sensor or (ii) via direct attachment of the thiolated oligonucleotide to the Au electrode surface. Both immobilization techniques were equally efficient and produced reproducible detection signals. The sensor was successfully tested on real samples and certified reference materials. The main advantages of QCM, such as low cost and high speed of analysis are hampered by the necessity of using PCR, which is both an expensive and slow method.163,64 A novel method for real-time determination of small molecule interactions with DNA has been developed.165,166 It involves measurement of the binding

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profiles of analytes to a modified quartz crystal sensor surface. After modification of the QCRS surface, DNA is attached to gold coated sensors via a bifunctional linker whereupon the surface is exposed to a range of analytes using a flow injection system. The resultant sensor is able to show reproducible changes in frequency which is dependent upon the binding modes of analytes. The magnitude of the frequency change and the binding profiles can be directly related to the strength of analyte-DNA binding. Using this method Fucassi et al. could differentiate between covalent bonding, intercalation and chemically induced DNA cleavage. The system has been used to study catalytic antioxidant behaviour and the fates of genotoxins from vitamin C reactions with lipid peroxides.165,166 A big problem with DNA/RNA based QCM sensors is their low detection limit which is in the range of 10ÿ7±10ÿ8 M.167 Such a low detection limit is probably due to difficulties with nucleic acid immobilization and low hybridization yields on the surface caused by steric hindrance of the hybridization reaction. Enhancement of the QCM response and lowering the detection limit in DNA sensing was achieved by increasing the surface area of the gold electrode.168 This was done by immobilizing 11 nm colloidal particles of gold on the gold electrode of the QCRS. Only 1,6-hexanedithiol containing two ±SH groups was efficient in immobilizing colloidal gold, whereas two other cross-linkers, cysteamine and cystamine containing only one ±SH group, yielded inconsistent frequency changes. The amount of the immobilized DNA containing ±SH group at the 50 -phosphate end was almost linearly proportional to the amount of the immobilized colloidal gold. Despite high theoretical sensitivity, low cost manufacturing and compactness that make real-time QCM biosensors attractive, there are many practical hurdles that have to be overcome.70 Both the physical properties of the liquid and the electrical short currents increase the electro-mechanical load on the resonator causing the oscillation quality to deteriorate. To achieve high practical signal stability and sensitivity, it is necessary to obtain a sufficiently stable resonance signal in the presence of even minute fluctuations in hydrostatic pressure due mainly to liquid pumping mechanics. Okahata and co-workers158,169 designed a flat flow-cell QCM chip for on-line analysis of l-scale samples. The use of 27 MHz quartz crystals instead of conventional 5±10 MHz improves the sensitivity of QCM sensors.169 A 27 MHz sensor signal had very low noise level (0.05 Hz) at flow rates up to 100 l/min. The 27 MHz QCM sensor was used to detect DNA-protein/peptide interactions in aqueous buffer solutions.169 From the time-dependence of the frequency shift, the binding amount, association constant, binding and dissociation rate constants were obtained. When DNA was immobilized on a QCM crystal, binding kinetics of antibiotics, a zinc finger-type peptide, a restriction enzyme, and a polymerase to DNA in aqueous buffer solution, were obtained. The 27 MHz QCM proved to be a highly

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sensitive and quantitative technique to detect various molecular interactions of ds-DNA in vitro.63,158,169±172 The electrochemical behaviour of phospholipid bilayers was investigated using the electrochemical quartz crystal nanobalance (EQCN) and quartz crystal immittance (QCI) techniques.173 The bilayer film of L--dipalmitoyl phosphatidylcholine (DPPC) formed on on Au electrode of the quartz crystal surface was modified with gramicidin (GR) forming ion channels permeable to Tl+ ions. The quartz crystal immittance measurements have revealed that the GR-modified DPPC membrane on the gold electrode behaves as a solid rigid film. Thallium is deposited at the interface Au-lipid film and forms an interlayer. GR ion channels incorporated in the membrane support Tl+ transport to the electrode surface underneath the lipid film. QCR sensors coated with lipid layers and embedded receptors have been used recently to study model membrane systems and their interactions with different ligands.174 It has been observed that the lipid layers formed in direct contact with the electrode surface do not fully mimic a real membrane because they may have a different structure; also the surface roughness of the sensor prevents an undisturbed organization of these layers. This problem has been overcome by making a sandwich structure with a soft polymer `cushion' between the sensor surface and the lipid layer.174±176 Using model membranes immobilized on the QCR sensor kinetic and thermodynamic parameters of ligand ± receptor interaction could be studied.77 The binding constant of gangliosides (ligand) with peanut agglutinin ± a 110 kDa lectin, was measured and this approach was further extended to studying the adsorption of bacterial toxins.60,77,177 To detect the target analyte, an enzyme-induced precipitate formation can be used. The immobilized or soluble enzyme converts the analyte into a precipitate that can be measured by QCM. Using this approach, detection of organophosphorus and carbamate pesticides and insecticides was performed.178-182 It should be noted that these measurements required an arrangement of a rather complex system comprising up to three different enzymes such as the one shown below.182 AChE

ChO

HRP

+ O2

+ DAB

acetylcholine ÿ! choline ÿ! H2O2 ÿ! 4,40 -diimino-3,30 -diaminobiphenyl (precipitate)

where: AChE is acetylcholinesterase, ChO ± choline oxidase, HRP ± horseradish peroxidase and DAB ± diaminobenzidine. Pesticides inhibit AChE thus causing increase in the concentration of acetylcholine, which is detected by QCM as a frequency shift due to precipitation of the product of DAB peroxidation. Interaction of humic acid semi-permeable membranes with glucose was investigated with the QCM technique. For this purpose, multi-layered films of humic acids were self-assembled in the presence of ferric ions on the quart sensor surface and their interaction with glucose was monitored. It was shown that membrane permeability to glucose could be regulated by the number of the

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layers. The information thus obtained can be used for constructing glucose sensors.183 Layer-by-layer self-assembly of complex multi-component films containing polymeric ion exchangers and enzymes urease (Ur), and arginase (Ar) was characterized with QCM.184 Poly(dimethyldiallyl ammonium chloride) and sodium poly(styrene sulphonate) were used as cation and anion exchanger respectively. Optimal multiplayer composition and architecture were determined and the membrane was used to make an L-arginine selective electrode. Awareness of the limitations of fluorescence and radioisotopic screening techniques has led to increased interest from pharmaceutical companies in labelfree screening techniques.185 Whilst techniques such as surface plasmon resonance (SPR) and thermal analysis are now commonplace with pharmaceutical analysis, these too have boundaries with respect to their application base, flexibility and sample requirements. QCM is a particularly useful labelfree technique for detecting drug-receptor interactions such as between low molecular sugars and their host receptors in liquid media in real time.185 Some other important bioanalytical applications of QCM such as study of the kinetics of protein adsorption or biofilm formation do not require the presence of an immobilized biomolecule.186 In such cases a surface of interest is either the quartz itself or a film of another material attached to the quartz crystal sensor. Protein adsorption is particularly relevant to the performance of biomedical devices and biomaterials as it is the first rapid event that occurs upon blood or tissue contact with their surfaces. Biocompatibility depends on a favourable interaction between the artificial surface and blood.187 Of significant importance to blood-surface interactions is the protein fibrinogen (Fg), which rapidly adsorbs to surfaces, polymerizes to form fibrin leading to a surface-induced clot and the fouling of biomaterials. Evans and Schoenfisch187 used the QCM method with resistance measurement capabilities and radiolabelling to probe the properties of Fg at various surfaces to characterize the surface activity and function of the adsorbed fibrinogen. Fg adsorption on unmodified and modified quartz depends on the shear rate suggesting that it is controlled by the diffusion factor.188 Interactions between protein layers were studied using QCM.189 A masssensitive linear frequency decay was observed with increasing thickness of the surface coating up to 20 protein layers. To engineer a controlled multi-layer coating, a combination of experimental techniques was used including selfassembly of thiolated peptides, covalent attachment of aminodextran hydrogel, and self-assembly of biotinylated proteins followed by attachment of polymerized streptavidin. The frequency shift
f of the sensor in phosphate buffer solution was about four times higher than
f of the coated sensor dried in argon, and this difference was explained by the presence of a significant amount of the entrapped water which accounted for 70% of the mass loading. The QCM data were in agreement with optical measurements. Protein films at the water/QCM interface can be quantified by measuring
f , but the adsorbed protein layers are not rigid and have some structural flexibility

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or viscoelasticity that is undetectable by simple frequency determination.190 Viscoelasticity can, however, be quantified as the energy loss, or dissipation of the shear movement of the crystal in water. Very small structural and orientation changes of an adsorbed protein layer, including chemical cross-linking, have been monitored by QCM-D with high accuracy. QCM with the dissipation technique has been shown to be a sensitive tool to study protein adsorption kinetics in aqueous solutions, with sensitivity in the ng/cm2 range.8,190,191 It provides information about mechanical and structural properties of the protein films such as viscoelasticity. This method also measures water associated with the hydration layer of proteins and/or polymer films.192 HoÈoÈk et al.193 used the QCM-D technique to study the effects of pH and ligand-binding on the kinetics of changes in f …t† and D upon loading of haemoglobin onto a hydrophobic methyl-terminated, thiolated gold surface. Simple monolayer formation of presumably intact proteins was observed at high ionic strength near their isoelectric points (pI), and bilayer formation around pI was found at lower ionic strength. Much lower saturation coverage at low and high pH compared with pH > pI, which was attributed to electrostatic repulsion due to surface charge on the proteins and increased spreading (conformational changes) because of reduced protein stability.193 QCM-D measurements at multiple harmonics were used to investigate a number of biomolecules in the adsorbed state.190,194 The QCM data were further supported by D2O substitution and ellipsometry measurements, which were used to quantify the amount of coupled water detected by the QCM. The quartz crystal sensor surface can be coated, in principle, with any biomolecule and utilized for the development of biofunctional patterned surfaces. This was done by extending lipid- and protein-based immobilization strategies to surfaces patterned with alternating areas of Au and SiO2. After a number of sequential immobilization steps, all of which were followed either in parallel or simultaneously using QCM-D, Au was functionalized by DNA, while SiO2 was covered by an inert phospholipid bilayer. Since the phospholipid bilayer is not only inert towards protein adsorption, but also towards lipid vesicle adsorption, this pattern allows controlled specific immobilization of DNAtagged lipid vesicles, which may or may not carry protein functionality.190,194 QCM has been used to study the fouling of metals exposed to protein solutions and the efficiency of surfactants in cleaning up metal surfaces.195 The QCM surface coated with gold or chromium was exposed to 3 wt% solutions of pure -lactoglobulin and a commercial skimmed milk powder at neutral pH, and before and after heat treatment at 80 ëC. The piezoelectric sensor was also used to evaluate the cleaning action of surfactant Tween 20 and enzyme trypsin. The mass load changes on the QCRS were interpreted in terms of the hydrodynamic thickness of the adsorbed films and led to the conclusion that QCM is a sensitive and convenient technique for monitoring detergent efficiency.195

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A recurrent problem in QCM applications is the reproducibility of the experimental results. Lipids and proteins can adsorb irreversibly and nonspecifically onto the quartz crystal surface making its regeneration for subsequent re-use a significant challenge.77,196 It has been suggested that proteases are used for surface cleaning from proteins or to use disposable sensors.196 QCM coated with a pH-sensitive amphoteric polymer has been used to monitor cell growth.197 The sensor responded to the biochemical products of cellular metabolism rather to cells, and it did not require immobilized biological receptors to monitor cell growth and metabolism. Microbial biofilm formation on metallic surfaces has been studied by a combination of EQCM with gravimetric and optical methods.198 The use of flow-cell EQCM allowed real-time monitoring of biofilm formation and resulting open circuit potential UR. Measuring UR was particularly useful for assessing different methods for surface cleaning from microbial biofilms. It was shown that surface treatment with 10% H2O2 completely restored the original UR value, whereas disinfection with 70% ethanol did not affect it.198 The quartz crystal resonant sensor was capable of measuring the adhesion of Staphylococcus epidermidis to fibronectin-coated surfaces. Changes in resonant frequency were recorded and showed a linear relationship with the logarithm of cell suspension concentrations ranging from 1  102 to 1  106 cfu/ml.199 A non-invasive, real-time QCM method for the monitoring of cellular integration within commercial collagen-based dermal replacement scaffolds was reported.188 An unexpectedly high degree of acoustic energy transfer through heavily hydrated thick films (up to 0.5 mm) of collagen/glycosaminoglycan scaffold materials intimately associated with a quartz crystal sensor allowed for quantitative resonant frequency measurements on application of fibroblast cell suspensions to the material. Changes in resonant frequency and energy dissipation were commensurate with cellular interaction with the gel. Many marine organisms attach to solid surfaces with an extraordinarily strong adhesion. QCM-D was used to study adhesives from two marine organisms, the common blue mussel Mytilus edulis, and the brown algae Laminaria digitata. The cross-linking ability of adhesive byssal proteins from M. edulis (Mefp-1) is based on the presence of the reactive amino acid 3,4dihydroxyphenylalanine (DOPA), which can be oxidatively crosslinked by the enzyme tyrosinase. The brown algae L. digitata contains polyphenolic biopolymers which can be oxidatively crosslinked with bromoperoxidase (BPO). Chemical cross-linking of both proteins can be done with NaIO4. Adsorption of Mefp-1 and PP was measured by a decrease in resonant frequency and increase in dissipation. Addition of NaIO4 resulted in a significant decrease in frequency and viscoelasticity of the adsorbed layers due to crosslinking, which was also confirmed by ellipsometry.200 Monitoring cell growth or adhesion using QCM seems to offer a significant advantage over biological methods in terms of complexity of the method and

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duration of the analysis, but like protein adsorption, it is doubtful that there is a simple linear correlation between the frequency shift and the mass of the cells attached to the crystal surface.157,201 The frequency shift calculated for a monolayer of cells attached to a 5 MHz QCRS using the Sauerbrey equation (13.1) was at least ten times higher than the experimentally observed
f .201 The most likely explanation of this discrepancy is that the cell layer has visco-elastic properties and energy dissipation should be taken into account. Even very thin (a few nm) biofilms dissipate a significant amount of energy owing to the QCM oscillation.202 Various mechanisms contribute to this energy dissipation. Three main contributions suggested by Rodahl et al.202 are: (i) viscoelasticity and porosity of the biofilm that is strained during oscillation, (ii) liquid entrapped between the quartz crystal surface and the biofilm layer, which can move between or in and out of the pores due to the deformation of the film and (iii) the load from the bulk liquid which increases the strain of the film. These mechanisms constitute an effective visco-elastic load hence, biofilms cannot be considered to be rigidly coupled to QCM oscillation and data interpretation of QCM measurements of biofilms requires more sophisticated models. A comprehensive model for `adhered cells-quartz surface' system has been suggested.203 It considers three layers, namely a rigid layer of the extracellular matrix, a visco-elastic layer of cells and a water layer between them. The QCM-D method used to study cell attachment and spreading over the polystyrene coated quartz crystal surface showed correlation between the degree of dissipation and amount of cells attached.204 An unusual QCM set-up allowed very sensitive detection of bacteriophage.205 By immobilizing ligands that interacted specifically with phage coat proteins on the QCRS surface, selectivity of phage adsorption was achieved; then the amplitude of the surface oscillation was then increased to `shake off' the adsorbed phage. The authors claimed that the quantification range spanned over at least five orders of magnitude and the method was sensitive enough to detect as few as twenty phages. Such sensitivity makes this method attractive for virus detection. Not only can QCM be used to detect cells, but living cells immobilized on the QCR sensor can also be used to study effect of chemical substances on their behaviour. The QCM technique measuring
f and
R detected microtubule (MT) alterations caused by the drug nocodazole in the immobilized endothelial cells (EC).206 Depolymerization of microtubules was dose dependent in the concentration range of nocodazole between 0.11 and 15 M. By relating QCM data to the microscopic examination of endothelial cells, it has been shown that different regions on
f ÿ t and
R ÿ t curves correspond to different events that change mass and viscoelasticity of the EC layer.207 These events reflect different stages of cell-surface interactions such as contacting of the surface, adhersion and spreading. In the examples quoted above the quartz crystal resonance sensor technique was either used as only a mass change detector, or as a mass and dissipation

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detector. It is also possible to design a QCR biosensor that exploits only the viscosity-density change of the liquid medium. In an experimental set-up the concentration of a biomolecule is measured in the solution by carrying out a sol-gel transformation which significantly influences the viscosity of a medium.208,209 The latter is detected by a quartz TSM sensor. If there is no film formation on the quartz crystal surface, the frequency and resistance changes of the sensor depend only on the liquid density and viscosity. This approach was realized to measure the concentration of fibrinogen in a clotting assay and endotoxin in the limulus ameobocyte lysate LAL test which is also based on the clotting of the LAL substrate and the endotoxin. It has been suggested that the Kanazawa±Gordon equation (13.4) and eqn 13.10 are applicable in these systems:210 r  `  ` 13:10
R ˆ 8K 2 C0 f0 q q where K 2 is the electromechanical coupling factor of the quartz and C0 is static quartz capacitance. A polystyrene spin-coated QCR with a fundamental frequency of 10 MHz was used to record frequency and resistance changes during the LAL-endotoxin clotting. Different approaches to produce a calibration plot were examined and it was found that the time to reach a certain resistance shift gave a linear graph `time vs logarithm of endotoxin concentration' at low endotoxin levels. At higher endotoxin concentrations, the
R or
f at a certain time gave a better correlation; the combination of both methods allowed fast measurement of endotoxin concentration in the range 100 pg/L to 2 g/L.210 QCM immunosensors Immunosensors are biosensors in which the immunochemical reaction is coupled to a transducer.211,212 The fundamental basis of all immunosensors is the specificity of the molecular recognition of antigens (Ag) by antibodies (Ab) to form a stable complex. This is similar to immunoassay methodology. Immunosensors can be categorized by the detection principle used. The main types are electrochemical, optical and microgravimetric immunosensors. In contrast to the enzyme-linked immunosorbent assay (ELISA), which is most commonly used to detect Ag-Ab interactions, modern transducer technology enables the label-free detection and quantification of the immune complex. Further potential advantages of immunosensors vs conventional immunoassays include lower cost and a shorter time of analysis. The analysis of trace substances in environmental science, and pharmaceutical and food industries is a challenge, since many of these applications demand a continuous or a nearcontinuous monitoring. The use of immunosensing sequences in these applications is a possibility. Similarly, continuous monitoring of certain analytes

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Surfaces and interfaces for biomaterials

can be used in clinical diagnostics. The future role of this technique in the laboratory as well as for bedside testing will become even more important as the clinical laboratory is faced with increasing pressure to contain costs.211 Shons et al.213 first applied antigen coated piezoelectric crystals to immunoassay. The first piezoelectric immunosensors lacked reproducibility and stability when immersed in a liquid, so a measurement technique known as `dip and dry' was used.213±215 According to this protocol, the QCR sensor (QCRS) with immobilized antibody was dipped into a solution containing antigen, rinsed, dried and
f was compared with that of the sensor before dipping. Using this technique, herbicide atrazine, Candida albicans and IgM (human) were detected. In one case atrazine was detected in the range from 0.001 to 1 ppb, which is a remarkable sensitivity.216 Nevertheless, dip-and-dry technique suffers from intrinsic problems ± it is slow and cannot be used in the continuous regime. More commonly QCM for immunosensing is used in the continuous mode in liquid phase.23,217 The key issue in achieving stability and reproducibility of analysis is the choice of the immobilization method of the immune reactant on the quartz crystal surface. As with biosensor fabrication, immobilization can be done in a variety of ways. Physical adsorption of antibodies is an attractive method due to its simplicity and it can be universally used to coat any surface.218 This is based on on irreversible adsorption of high molecular mass protein molecules.219 Electrical impedance analysis has been used to study Ab-Ag interaction on a polystyrene-coated quartz crystal surface.220 The motional resistance, R increases during both the Ab immobilization and the subsequent Ab-Ag binding indicating more power dissipation in the system. Resonant frequency, f decreases in both cases reflecting increased mass loading.
R in the Ab-Ag reaction is considerably larger than that in antibody immobilization process while
f are similar in these two processes. The ratio
R/
f therefore increases three-fold for Ab-Ag interaction vs. Ab immobilization.220 Physical adsorption is a statistical process driven by van der Waals forces, and adsorbed molecules have an irregular orientation; moreover, the adsorbed proteins undergo conformational changes that affect their properties such as bioactivity.221 Using AFM, it was shown that non-covalent attachment of antibodies ± sheep or mouse IgG ± to the gold coated mica resulted in large aggregated structures comprising randomly oriented antibodies.218 Not surprisingly, such a random orientation resulted in a lower efficiency of AgAb binding as measured by QCM. Another complication is the Vroman effect, that is, displacement of one set of adsorbed proteins by other proteins.222 These significant drawbacks of immobilization by physical adsorption make it a less popular technique for analysis in the liquid phase than alternative methods. Problems with low stability and reproducibility of sensors with physically adsorbed antibodies led to development of alternative approaches for binding of

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immune reactants to the QCR sensor surface. The Au electrode surface does not have surface sites for a strong attachment of biomolecules; to create these, several techniques have been used. They include the formation of: (i) a polymer layer with functional groups; (ii) a molecularly imprinted polymer layer; (iii) a self-assembled monolayer with functional group, (iv) an adsorbed layer of molecules with high affinity towards the immune reactant. In the latter case, formation of the interacting couples `protein A-immunoglobulin G', `streptavidin-biotin' or `avidin-biotin' is most commonly used. The initial step in the immobilization procedure is the preparation of bare QCRS electrode surface. In most cases the electrode is made from gold ± an inert material whose properties do not change in time. Its surface, however, may get contaminated due to inevitable adsorption of volatile organics and water upon storage.223 Although cleaning of the electrode surface has not always been used prior to immobilization,224±228 it has been shown that the frequency of cleaned QCM stabilizes faster that that of non-cleaned QCM,229 and most researchers carry out cleaning of the electrode surface before its modification.215,217,230±233 Pre-treatment of the QCR sensor in a glow discharge plasma chamber is used as a method of its surface cleaning.234 The majority of protocols for surface cleaning include treatment with a strong alkali solution (usually NaOH) followed by rinsing and treatment with a strong acid (HCl).235,236 Alternatively, the gold electrode surface is cleaned using `Piranha' solution (one part of 30% H2O2 in three parts of concentrated H2SO4).218,237,238 Polymer coating of QCM electrodes is commonly used to derivatize their surface, through which biomolecules are attached.239 Polymers can be immobilized by dip-coating,231,233,235 in which case the polymer is usually spread from a non-aqueous solution on the electrode and then air-dried. Polymer coating by adsorption from aqueous solution is based on hydrophobic interaction between the electrode surface and polymer.229 Polyethylenimine is often used for this purpose, as it has reactive amine groups through which immune reactant molecules can be attached. Koning and GraÈtzel231 elaborated a sensor (coating with polyethylenimine followed by glutaraldehyde, prior to the attachment of antiglycophorin A antibody) for human erythrocytes. Coating by chemical attachment of ( -aminopropyl)triethoxysilane ± (silanization) is also used to create reactive amine groups on the surface.233 For the immobilization of biological materials on the electrode surface, electropolymerization has some advantages, because this process is easier to control and films have a more uniform thickness and chemical composition and are mechanically and chemically stable.240,241 Glow discharge plasma polymerization has been used to make thin polymer films from ethylenediamine and allylamine, which also contain amine functionality.234,242,243 Further immobilization of antibodies, antigens or other biomolecules is accomplished using adsorption or a chemical coupling agent such as glutaraldehyde.229,234,235 Immobilization of antibodies

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Surfaces and interfaces for biomaterials

via aldehyde groups of glutaraldehyde on a quartz crystal precoated with PEI produces a layer of randomly oriented antibodies.235 Considering the high costs of QCM sensors as the main obstacle towards commercialization, Su suggested use of silver electrodes instead of gold.244 Silver is less inert and less stable than gold and a polystyrene coating was used to protect the silver electrode from undesired oxidation and immobilize biomolecules to make an immunosensor.196,245 Lin et al.246 have argued that molecularly imprinted polymer layers (MIP) formed on the quartz crystal surface in the presence of analyte have affinity and selectivity comparable to that of natural receptors, and they are easier to prepare than Ab-based sensors. The absence of biomolecules on a QCRS ensures its higher physico-chemical, mechanical and thermal stability. A mixture of 3-dimethylaminopropyl methacrylamide with different acrylate crosslinkers and albumin was copolymerized and spin-coated on a Au-coated quartz crystal surface. Albumin was extracted from the MIP with 20% methanol in water. Thus formed, the MIPQCM showed higher selectivity towards albumin vs other proteins of a lower molecular mass using model mixtures and blood serum. The MIP sensor also had faster response times and greater adsorption capacity towards analyte. The covalent coupling of the Fab' fragment of antibodies via its free thiol groups to a monolayer of phosphatidylcholine and cholesterol on QCM was used to create an immunosensor with high antigen binding efficiency.247 Protein A, a cell surface receptor produced by Staphylococcus aureus, has been widely used for the quartz crystal surface coating in immunosensor design. This protein has high natural affinity with the Fc region of immunoglobulins, particularly IgG molecules.248 Despite the absence of chemical interaction between protein A and IgG, the dissociation constant KD for the protein A-IgG complex is very low; it varies between 10ÿ9 M and 10ÿ6 M assuming pseudofirst-order of this process.249 For comparison, highly specific Ab-Ag interactions form complexes with KD values around 10ÿ9 M.250 Such a high affinity between protein A and immunoglobulin provides stable binding of IgG to QCRS surface coated with protein A. One protein A molecule can bind at least two IgG molecules.251,252 Protein A is a single polypeptide chain of molecular weight 42,000.251 If physically adsorbed on a QCRS surface without losing its ability to bind IgG molecules and without blocking the active sites of the antibodies for binding target analyte.218,253 Although protein A adsorption is of a physical nature, driven by van der Waals forces, it is irreversible like any other protein adsorption, and the gold-protein A complex is highly stable.217 Guilbault et al.215 utilized polyclonal antibody attached to the gold electrodes pre-coated by protein A. A big advantage of using protein A coating is that antibodies can be directly attached to it and no additional surface activation is required,218,235 whereas immobilization of IgG molecules on other polymer films requires use of a coupling reagent that activates surface functional groups.

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Attachment of antibodies to the surface via protein A has the advantage of preferential orientation of the immobilized molecules, as their Fab domains are likely to be exposed to the external medium.217,231,254 Use of protein A for immobilization of antibodies produced better stability and reproducibility of results in QCM detection of Salmonella typhimurium compared to the QCM sensor designed by covalent attachment of antibodies to the PEI film via glutaraldehyde activation. The sensitivity of both sensors was comparable.235 Like antigen-antibody interaction, streptavidin (avidin)-biotin interaction is highly specific and irreversible, with KD values reported as low as 10ÿ16 M.241 It has been used for the immobilization of biomolecules in a QCM immunosensor. Dupont-Filliard et al. coupled biomolecules to the QCM surface coated with polypyrrole-biotin film through an intermediate avidin layer. The procedure involved three steps: (i) electrosynthesis of the polypyrrole-biotin film at an electrode surface; (ii) surface immobilization of the avidin layer via the biotin entities and (iii) anchoring of the biotin-labeled biomolecules on the polypyrrole biotin/avidin layer to create a bioactive surface. The authors described several advantages of using this technique, namely, (i) avoiding the use of chemical reagents that may affect biomolecules, (ii) high precision of immobilization of biomolecules only at the surface of the film, which leads to a high sensitivity, (iii) versatility of the technique, as a wide variety of commercially available biotin conjugates is available. A DNA sensor was prepared by immobilization of biotinylated DNA probes creating the sensing layer, polypyrrole-biotin/avidin/ DNA probe. The sensor was regenerated by destroying the biotin/avidin complex using a detergent solution via the solubilization of the avidin layer. Removal of the avidin layer immobilized on a polypyrrole-biotin film did not alter the functionality of the latter. After the initial 15±20% loss of activity, regenerated biosensors exhibited a stable and reproducible response up to the 10th regeneration. Use of self-assembled monolayers in the QCM immunosensor design is based on the high affinity of thiolated compounds with a gold surface.255 The SAM technique is one of the simplest ways to produce a well-ordered layer suitable for further modification with antibodies. Based on the immobilization of antibodies onto a SAM of cystamine with sulfo-succinylmidyl 4-(pmaleimidophenyl)butyrate (sulfo-SMPB) as linker, a piezoelectric immunosensor was fabricated for sensing C. trachomatis with a detection limit of ~260 ng/ml.256 Thiols are strongly attached to gold surface via chemisorption rather than physical adsorption forces. Hydroxy-, amino- and carboxylterminated thiols are commonly used for covalent immobilization of bioligands to SAM. A great variety of terminal functionalities in monolayers offers high flexibility in biosensor design using SAMs. Zhao et al. 257 coupled hydroxyl and amino-terminated SAMs to the 50 phosphate-end of DNA, while carboxyl-terminated ones were coupled to the 30 hydroxy end of the nucleic acid using a water soluble carbodiimide. X-ray photoelectron spectroscopy (XPS) and cyclic voltammetry (CV) data proved

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Surfaces and interfaces for biomaterials

that hydroxyl-terminated SAMs were more efficient for the covalent attachment of DNA than the other two types of SAM. A multi-layer QCM for Chlamydia trachomatis was designed by assembling a primary layer of cystamine on the gold electrode; the cystamine monolayer was further modified with sulfosuccinylimidyl 4-(p-maleimidophenyl)butyrate and the fragmented F(ab')(2) anti-mouse IgG Ab, to which the anti-C. trachomatis LPS-Ab was attached. The latter molecule was used for Chlamydia trachomatis detection. The sensor sensitivity for this bacterium was 260 ng mLÿ1 in urine and it showed long-term stability upon storage at 4 ëC. 256 Similarly, a piezoelectric immunosensor based on a self-assembled monolayer of cystamine has been developed for the determination of Schistosoma japonicum Ab in rabbit serum. Immobilization of the Schistosoma japonicum Ag to the positively charged SAM was achieved electrostatically using a negatively charged polystyrene sulphonate layer. It was shown that Ag immobilized by this procedure had higher immunological activity and binding efficiency than Ag immobilized with glutaraldehyde. The immunosensor had satisfactory sensitivity and detection limit.258 To increase the mass load on a QCM immunosensor, an increase in surface area for the Au electrode was suggested by coating it with porous gold. A 16fold increase of surface area was achieved, and this led to a 11.4-fold increase in thiol adsorption and a 3.3-fold increase in protein adsorption, thus amplifying the sensitivity of the original flat non-porous device.259 A piezoelectric immunosensor was developed for the rapid detection of Escherichia coli O157:H7. Antibodies were immobilized onto a monolayer of 16-mercaptohexadecanoic acid self-assembled on an Au electrode at an AT-cut quartz crystal. Covalent attachment of Ab to the carboxylic groups of the SAM was carried out with N-hydroxysuccinimide ester as a reactive intermediate. The immunosensor could detect the target bacteria in a range of 103±108 CFU/ml within 30±50 min. The proposed sensor was comparable to Protein A-based piezoelectric immunosensor in terms of the amount of immobilized antibodies and detection sensitivity.260 Similarly, a batch-type QCM system for detecting chloramphenicol (CAP) was developed. Anti-CAT Ab were covalently attached to the carboxylic groups of SAM of 3-mercaptopropionic acid via activation with water-soluble carbodiimide (EDC) and N-hydroxysulphosuccinimide. The antibody-immobilized sensor showed 10±50-fold enhanced sensitivity in comparison with the uncoated sensor or coated only with 3-mercaptopropionic acid. Repeated use of the sensor up to eight times was possible after 1 min regeneration with 0.1 M NaOH.236 These results proved that the QCM sensing has potential as a rapid screening method for low molecular mass compounds such as environmental endocrine disruptors, which usually require timeconsuming, complex and expensive labelling analytical techniques. The SAM technique can be used to immobilize biomolecules by labelling target bioligands with thiol groups. Thus, biomolecules become capable of selfassemblying on a QCM surface. Using this approach, thiolated DNA and

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oligonucleotides were used to produce DNA hybridization sensors.261 Immobilization of thiolated Ab produced variable results, perhaps due to the oxidation of ±SH groups.218 Molecules with disulphide links were suggested, e.g. DNA±(CH2)3±S±S±(CH2)3±OH, to minimize the undesirable side reaction of ±SH oxidation. The disulphide link can then be cleaved by reduction with dithiotreitol prior to immobilization. Thiolated anti-Salmonella antibodies to Salmonella typhimurium were directly attached to the Au electrode of a QCM sensor from a solution of Ab, and a heterobifunctional thiolation cross-linker, sulphosuccinimidyl 6-[3-(2pyridyldithio)propionamido]hexanoate.262,263 The immunosensor could detect 9.9  105 to 1.8  108 cells/ml within 30±90 min. For repeated use of the immunosensor, the efficiency of various regeneration reagents in removing adsorbed Salmonella cells and maintaining sensor sensitivity were compared. The best results were achieved with the use of 1.2 M NaOH. After the fifth assay, the decrease in frequency was only 24.4%, whereas treatment with 8 M urea solution led to the decrease in a frequency of 43.6% and treatment with 0.2 M glycineHCl, pH 2.8 resulted in a frequency shift by 63.2%. Rickert et al. also reported that regeneration of the QCM immunosensor with a 6 M urea solution. After this experiment, the sensitive layer was regenerated by destroying the antigen-antibody binding with a 6 M urea resulting in a 30% loss of sensitivity after three regeneration cycles.264 Several papers describe immunoassays in which the QCM sensor measures a frequency shift caused by agglutination of immunized latex microbeads. The method was named by its authors latex piezoelectric immunoassay, or LPEIA.265±268 A suspension of latex microbeads (0.2 m in diameter) with immobilized antibodies against the analyte reacts with antigen, as a result of the interaction, the latex particles agglutinate. An uncoated QCM registers an increase of mass as the oscillation frequency shift due to the precipitation/ adsorption of the agglutinated particles. LPEIA was used to measure clinically relevant fibrin degradation products (FDP) and C-reactive protein (CRP) in blood serum. Response of the QCM sensor to FDP in human serum was within 10 min and stabilized within 60 min. The frequency shift of the QCM sensor and the absorbance change at 570 nm of the photometric method showed good correlation. As the piezoelectric sensor is used uncoated, its regeneration can be carried out with piranha solution. The sensor was used at least three times without any loss of sensitivity and results were reproducible.269 A commercial prototype sensor for CRP detection has been designed. Similarly, changes in viscosity/density of solution resulting from an antigenantibody agglutination reaction were followed by the QCM technique.270 Using this approach, Staph. epidermidis was detected in the infected human serum samples. Use of QCM immunosensors is not limited to the detection of immune reactants or biomolecules. Environmental pollution resulting from dioxins is

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becoming huge problem in the world. A piezoelectric immunoassay was developed for the rapid detection of polychlorinated dibenzo-p-dioxins (PCDDs). The system uses a competitive inhibition enzyme immunoassay (EIA) based on a mouse monoclonal Ab against 2,3,7,8-tetrachlorodibenzo-pdioxin (TCDD) and a conjugate of a dioxin-like competitor coupled to the enzyme horseradish peroxidase (HRP). The anti-dioxin Ab was immobilized on a 10 MHz AT-cut quartz crystal resonator modified with a SAM of dithiobis-Nsuccinimidyl propionate. TCDD at different concentrations in the range 0.001± 10 ng mLÿ1 was mixed with a constant amount of HRP-conjugated competitor and the frequency response due to the adsorption of the samples on the biosensor surface measured. The results showed that TCDD could be quantitatively detected in the concentration range 0.01±1.3 ng mLÿ1. The sensitivity and selectivity of the QCM immunosensor is comparable to EIA and ELISA methods in the detection of PCDDs. The developed QCM immunosensing system offers significant improvements in speed, sample throughout and cost for the qualitative and quantitative detection of PCDDs compared with GC-MS.271 Regeneration of the modified sensor interface by dissociating the bound analyte from the antibody-coated sensor indicates the reusability of this biosensor. In this study, both 20 mM glycineHCl buffer (pH 2.5) and 6 M urea were used to regenerate the QCM biosensor. The results indicate that the sensing interface still retained more than 86% activity after four regeneration cycles performed with glycine±HCl buffer (pH 2.5). The recovered activity using 6 M urea was 70% after four regeneration cycles. To detect the herbicide atrazine, monoclonal Ab against atrazine were immobilized on a quartz crystal sensor via Protein A covalently attached to the gold electrode surface activated with 3,3-dithiobis(propionic acid) Nhydroxysuccinimide ester. The MAb±Protein A complex was stabilized by cross-linking with dimethyl pimelimidate. The immunosensor thus developed was able to specifically respond to atrazine with a linear response of the frequency shift to atrazine in the concentration range 1±200 g/mL, however, sensor regeneration and re-use were a problem.272 Anti-PCB polyclonal sheep Ab were immobilized on quartz crystal sensor via protein A. The QCM sensor thus obtained was successfully used to detect 4,40 DCB (dichlorobiphenyl) and 2,4,40 -TCB (trichlorobiphenyl); the sensor detected TCB in the concentration range 0.2±2000 g/L and, remarkably, it worked in non-aqueous solutions as well.273 Cooper et al. suggested a novel approach for using QCM immunosensing techniques.274 Instead of measuring the frequency shift caused by the interaction between type 1 herpes simplex virus (HSV1) and specific antibodies covalently attached to the QCM surface, the authors monitored detachment of the virions from the surface by monotonously increasing the amplitude of oscillation of the piezoelectric sensor, while using it to detect the acoustic noise produced when the interactions were broken. The method termed rupture event scanning

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(REVS) is quantitative over at least six orders of magnitude of concentration range, and its sensitivity approaches detection of a single virus particle (0.5 femtogram) in phosphate buffer solution containing 1 mg/mL BSA and 50 virions in serum. Sensitivity of REVS is several orders of magnitude better than that of QCM or SPR. The REVS technique has been successfully used to detect E. coli, N. meningitidis, adenovirus type 5 and S. aureus. Scan time is as short as 1 min and no amplification is required for detection.

13.4 Combination of QCM with other techniques 13.4.1 QCM and SPR Surface plasmon resonance (SPR) is a highly sensitive method currently used for monitoring processes occurring at or near interfaces.275 It is an optical method based on total internal reflection from an interface between a material with higher refractive index such as metal film on glass and a medium with lower refractive index such as liquid. The generated evanescent wave moves away from the interface into the medium penetrating it to a distance of around 300± 600 nm. Analysis of the evanescent wave profile gives information about the adsorbed layer and allows measuring its thickness and amount of the substance adsorbed.276 SPR has been used as a gas phase sensor and as a method of monitoring biochemical and cellular interactions in a liquid.277±279 Like QCM, it is a non-invasive, non-labelling technique that provides kinetic data in real time. SPR is often combined and compared with QCM.238,280 QCM measurements are more sensitive to the environment, such as high ionic strength solutions, drifts in temperature or mechanical disturbance that cause changes in solution viscosity and density. With tight temperature control, SPR measurements have better reproducibility and reliability. Careful solution preparation and temperature control improve QCM performance. On the other hand, the QCM instrumental setup is relatively simple and fast. Both methods are easily capable of monitoring protein deposition with high sensitivity, however, the resonance angle shifts observed in SPR when studying bacterial suspensions are very low. In the same system QCM was capable of measuring bacterial concentration as low as 1000 cells/ml although the reliability of the method has to be improved.238 In another system studied, interaction of glucose with its specific enzyme glucose-6-phosphate dehydrogenase (GDH), SPR failed to register the enzyme-substrate binding, whereas the QCM flow system clearly showed strong interaction between GDH and glucose. Apparently change of the refractive index was too small in this system.185,237 Because QCM and SPR monitor changes in different parameters it is not surprising that for certain applications one technique is more sensitive than the other. An important advantage of QCM over SPR may be a relative simplicity of surface modification of the piezoelectric crystal surface. This approach was used

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to study complement activation by different biomaterial surfaces such as titanium, polystyrene or poly(urethane urea).281 The sensitivity of both methods was comparable although it depended on the analytical cell design.

13.4.2 QCM and other techniques EQCM is the most obvious example of QCM coupled with another technique in the same instrument and its applications are discussed in other paragraphs of this chapter. In most cases QCM and other methods are used complementarily. Hess et al. used the scanning electrochemical microscope (SECM) in combination with EQCM to investigate standing acoustic waves over the quartz crystal.282 The adsorption kinetics of three human blood proteins ± human serum albumin, fibrinogen and haemoglobin from model solutions onto glass and quartz coated with 12 nm thick titanium dioxide film was measured using three different experimental techniques: optical waveguide light spectroscopy (OWLS), ellipsometry (ELM) and QCM-D.190 All three techniques proved to be suitable to study kinetics of protein-surface interactions in situ. They were also consistent in recording the adsorption kinetics trends and the results were used to quantify the adsorbed mass. Using different protocols of surface treatment, it was shown that protein adsoption was largely irreversible, and so was the adsorption of antibodies against these proteins. The two optical techniques produced comparable results in most cases, which could be converted into adsorbed protein (`dry') mass. Data obtained with QCM-D, on the other hand, differed significantly in terms of the adsorbed mass, being 1.75 to 3.2 times higher. The higher mass value calculated from
f includes both protein mass and water associated with the protein layer. Analysis of the energy dissipation in the adsorbed layer and its magnitude in relation to
f provides information about the mechanical and structural properties, such as viscoelasticity and shear modulus, of the adsorbed film and the surrounding bulk liquid environment.190,202,283 The combination of optical techniques and QCM revealed much more information than any of these methods used separately. By measuring the damping of the amplitude of the crystal vibration after the alternating electric field has been turned off, one can qualitatively determine how much energy the adsorbed layer can dissipate. Since the signal from rigid adlayers is only weakly damped, it can be said to have a small dissipation factor. In other words, there are relatively few degrees of freedom, into which energy can be lost. Conversely, soft adlayers have many degrees of motional freedom and a correspondingly large dissipation factor. The amplitude of the signal from a crystal with such a layer adsorbed on it decays very rapidly. The QCM-D instrument serves as a complement to optical waveguide lightmode spectroscopy. Both techniques are capable of low-background, real time, in situ measurements of ng/cm2 levels of adsorption of unlabeled samples from aqueous solution. However, they provide different information since OWLS can

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only detect the mass of the adsorbed molecule itself, whereas QCM is also sensitive to entrapped or closely coupled solvent molecules such as water. In addition, visco-elastic information inherent in the value of the dissipation factor has no equivalent in OWLS. The QCM-D has the further advantage that it does not require transparent substrates, allowing examination of substrates such as gold in addition to the metal oxides used in OWLS experiments.190,202,283 It has been pointed out that QCM is one of the few methods that allows precise measurement of the quantity of an ultrathin layer of an element deposited onto a solid surface.284 Other methods such as Auger or X-ray photoelectron spectroscopies provide only relative and semi-quantitative information, whereas nuclear or ion-scattering spectroscopies are limited to certain combinations of the adsorbate and substrate. Unlike QCM, all these methods require expensive equipment and complicated experimental conditions. Results obtained with inductively-coupled-plasma mass spectroscopy, used to determine the coverage of an ultrathin layer of antimony on gold, were consistent with QCM measurements on the same sample.284 Coupling of the tapping-mode atomic force microscope (AFM) and QCM has an advantage of obtaining simultaneous information at nano- and sub-nano(AFM) and micro- (QCM) levels.285,286 The AFM/QCM coupled technique proved useful for investigating a wide range of processes, from protein adsorption on metallic surfaces285 to metal electrodeposition.286 Copper and silver deposition on gold produces rigid films, whereas protein adsorption is dominated by viscous interactions. Respectively, the frequency variations of the overtone number n evolve as
fn =n for rigid films, while in the case of viscous interactions in which the thickness of the liquid interacting with the resonator is a function of the oscillation frequency, the frequency variations of the overtone number n evolve as
fn =n0:5 . Measuring overtones thus provides useful quantitative and qualitative information on the nature of the adsorbed film. A coupled quartz crystal microbalance/heat conduction calorimeter (QCM/ HCC) has been designed for very sensitive and simultaneous mass and heat flow measurements of thin films interaction with a solvent vapor in a controlled gas phase.287

13.5 Acoustic/piezoelectric sensors Acoustic/piezoelectric sensors are based on a piezoelectric substrate through which acoustic waves propagate. QCM operates with a bulk acoustic wave (BAW) and its mass sensitivity is limited by the crystal thickness.288 The surface acoustic wave (SAW) devices can operate using different modes of wave propagation.22,289±291 Rayleigh SAW sensors are suitable for gas media but not for liquids in which they suffer unacceptably high energy dissipation and loss of sensitivity because of the surface-normal particle displacement.67,292 In contrast, shear-horizontal polarized surface acoustic wave modes (SH-SAW) or

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plate modes (SH-APM) are not significantly affected by immersion in liquid because the particle displacement in these modes is horizontal ± parallel to the sensor surface. Currently they are becoming a subject of intensive studies as biosensors.293,294 SH sensors are less sensitive than the Rayleigh SAW sensors, so to increase their sensitivity the waves are propagated through a waveguiding layer.295,296 First SH-APM sensors showed poor sensitivity that has not been significantly improved.297 Among such devices those using Love waves have recently attracted particular attention due to their sensitivity.* Love waves are elastic SH waves that are generated in the substrate and guided by a thin film deposited on it.298 Their propagation is sensitive to mass loading and viscoelastic effects of the medium. Usually two interdigital transducers (IDT) ± transmitting and receiving ± are used. Love-wave devices are based on ST-cut quartz crystal as the substrate with a guiding layer made from a polymer such as PMMA,294 deposited or etched SiO2,299,300 or a bi-layer structure comprising silica and PMMA.292 The PMMA layer is more sensitive than the silica layer but the latter is chemically inert, so their combination in a bi-layer has some advantage over individual layers.292 The mass sensitivity, stability and temperature coefficient of frequency of the bi-layer Love-mode quartz acoustic sensor is better than those of PMMA- or SiO2-coated devices but the insertion loss in water remains significant. It has been estimated that Love-wave devices should be sensitive enough to detect a uniform monolayer of most substances.301 A typical value of operating frequency of Love wave devices is about 100 MHz, much larger than typical values of QCM devices, near 10 MHz.302 High sensitivity of 500 pg/ml was reported for IgG adsorption to a Love mode mass sensor with the guiding ZnO layer (4 m thickness) on ST-cut quartz substrate. To increase selectivity of adsorption, an additional film of Au (20 nm thick)/Cr (10 nm thick) was deposited on the zinc oxide layer.303 ST-cut quartz has a low dielectric constant compared to water, hence being suitable for less polar liquids, its small piezoelectric coupling coefficient makes it less efficient in aqueous solutions.67 An alternative substrate for SH-SAW sensor is based on a 36ë rotated Y-cut X-propagating LiTaO 3 . 304 It has high dielectric constant and high electromechanical coupling coefficient. With a protective layer of SiO2 (1 m thickness) and an additional gold top layer (50 nm thickness) a SAW immunosensor was designed. It was coated with protein A to which rabbit anti-mouse IgG F(ab)2 fragment (antibody) was immobilized and subsequently the sensor was used to detect the goat anti-rabbit IgG (antigen). The sensor showed expected selectivity but further improvement of the sensor design would be required to improve its sensitivity.67 An in vitro assay to monitor the specific binding of an ultrasound contrast agent to a receptor immobilized on Love-wave biosensor has been developed.305 An antibody was coupled to the sensor via the * They are named after the English mathematician AEH Love, who described them in 1911.

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carboxylic groups of the monolayer of 11-mercaptoundecanoic acid selfassembled on the surface. Love-wave immunosensors with immobilized AM13-PVIII antibodies against a major phage coat protein PVIII have been successfully tested in detection of M13 bacteriophage.302 The antibodies were immobilized on the surface using dithiobissuccinimidyl-propionate. For the lowest bacteriophage concentration tested (1.6  109 pfu mlÿ1) the relative frequency shift was between ÿ14 ppm at 10 min and ÿ69 ppm at 2 h, which was better than ÿ7 ppm reported for a QCM based sensor.306 The response time was relatively slow for high concentrations of bacteriophages ÿ2h, but better than is usually required for routine biological analysis. In commercially available SAW transducers interdigital structures are usually made from aluminium or chromium by photolithographic technique. These electrodes corrode even in pH neutral buffer solutions, so that surface coating of SAW sensors is also necessary to protect the electrodes against corrosion.307 Protective polymeric films can also be used for covalent immobilization of biomolecules.308 It has been suggested that for covalent immobilization of biomolecules to SAW sensor surface, milder conditions and reagents should be used than for fabricating QCM biosensors because commercially available lithium tantalate SAW devices contain aluminium structures. Examples of immobilization methods suitable for engineering SAW biosensors include BrCN activation of the polyimide film,309 silanization of silica coating310 and photoimmobilization.308 Photoimmobilization requires derivatization of protein molecules with photoreactive groups. To produce sensing layers, proteins were modified at the lysine group with 3-fluoromethyl-3-(m-isothiocyanophenyl)diazirine (TRIMID). Upon light activation, covalent coupling of the protein molecule to the polyimide coating is achieved.308 Similarly, a dextran layer was attached to the surface. Dextran has a versatile chemistry allowing further covalent attachment of biomolecules.311,312 BrCN activation technique and immobilization of anti-mouse IgG (antibody) led to the design of a mass immunosensor with sensitivity of 58 Hz/pg by mouse IgG (antigen) and 51 Hz/ pg for the immobilized anti-glucose oxidase (antibody) by glucose oxidase (antigen)309 but the whole immobilization process took a week to complete and reproducibility of the sensor was unsatisfactory. Aluminium structures corroded in the course of the immobilization. On the contrary, photoimmobilization of proteins proved to be reproducible but it produced lower sensitivity of detection and still was quite complex a procedure.308

13.6 Future development of piezoelectric sensors There are general trends in the development and analytical applications of acoustic/piezoelectric sensor techniques in liquid media, which should further improve their performance:

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· improvement of sensors and whole electronic circuits; · consideration of energy dissipation, impedance, conductivity, surface nonuniformity and roughness, stress and compressional waves using comprehensive theoretical and numerical models and development of relevant software; · application of piezoelectric sensors in combination with other techniques such as spectroscopy, microscopy and chromatography; · use of sensor arrays for analysis of complex biosystems; · use of new piezoelectric materials such as GaPO4 may become more widely applied; · development of new setups possessing greater sensitivity than conventional QCM method; · miniaturization of sensors and cells; · increase in working frequency of QCM crystals to 30 MHz and higher; · improved stability and reproducibility of other acoustic/piezoelectric sensors; · direct measurements of adsorption of small molecules specifically interacting with surface functionalities; · development of new experimental setups for real-time and on-line use of piezoelectric sensors; · new more specific methods of surface modification and immobilization of bioreceptors; · immunoassays in organic solvents convenient for investigations of hydrophobic molecules; · investigation of bioaffinity interactions on the interfaces using advanced methods (impedance scans, REVS, etc.); · development of new biosensors using a variety of biofunctionalities with protein, e.g., RNA, DNA, lipids, saccharides, enzymes.

13.7 Thermal gravimetry A thermal analysis (TA) technique enables efficient measurement of changes of physical parameters such as mass (TG, DTG), temperature (DTA), enthalpy (DSC), together with various physical properties determined through mechanical (TMA), optical, magnetic, electrical, acoustic, emanation and other observations. These measurements take place during heating of materials using a controlled temperature program. Evolved gas analysis (EGA) may be based on chemical measurements, depending on the method of detection used. Consequently, an essential subsequent interpretation step is usually required to relate these physical TA measurements to each chemical change that is of interest.313 Available instruments permit a rapid and efficient completion of thermal experiments, obtaining quantitative data that are suitable for measuring weight loss upon heating, reaction enthalpy, specific heat, reaction temperature, melting point and reaction rate. Thermal changes, including chemical reactions,

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are often more complex than recognized, so that kinetic data and their interpretation may be incorrect. In measuring rates of thermal changes, the following factors may influence kinetic characteristics: reaction reversibility, melting, reactant self-cooling/heating, multiplicity of rate processes (concurrent and/or consecutive), and others.313 The term TKA (ThermoKinetic Analysis) describes any research in which physical measurements are usually carried out to determine the course and/or extent of the changes occurring in a sample held at constant or programmed temperature environment. It allows one to elucidate the reaction mechanism (in the usual chemical meaning of this term) and identify parameters that determine absolute and relative levels of reactant reactivity and control reaction rates. These objectives are generally approached through experimental investigations which attempt to find out what chemical reactions have occurred (stoichiometry), how fast (kinetics) and obtain any other information capable of elucidating all relevant aspects of the changes identified such as textures from microscopy, structures and topotaxy from crystallography, and bonding from spectra.313,314 Enthalpy changes can contain composite contributions, for example, where there are two (or more) concurrent chemical reactions and/or a phase change, such as progressive melting and/or reactant sublimation. Some types of change in the reacting system are not detected by particular TA methods, for example, TG does not record enthalpy changes due to a reaction, melting or other crystallographic transformations. To overcome method-specific limitations, complementary TA techniques may be used; TG and/or DTG are often accompanied by DSC and EGA with chemical and/or physical detection methods. Another limitation is that TA techniques, in contrast with microscopic observations, do not give direct evidence of the changing (geometric) dispositions within the reactant particles. Reaction geometry is often inferred indirectly through the interpretation of reaction rates by heterogeneous, interface advance, kinetic models.313±315 There have been many publications concerned with general and specific aspects of the experimental methods and theory of TKA, 316 together with specialist surveys and theoretical discussions.313±315 These citations refer to an extensive, now predominantly older literature, which has been recently revised.316,317 Current concepts of crystal chemistry, that form the theoretical foundation for TKA kinetic analyses, are incapable of constituting a comprehensive culture for continuing investigations of the variety of thermal changes that occur on heating reactants in condensed phases.313 Thermal gravimetry is a destructive method, which has lower sensitivity than QCM and has limited applications in investigation of interfacial phenomena in liquid phase.

13.8 Non-QCM adsorption methods Measuring of the adsorbed amount of gas and recording the adsorption isotherms are widely used for analysis of the surface structure of materials.318,319

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These measurements can be performed by gravimetric or volumetric methods. Conventional methods for the adsorption of gases and vapours such as N2, Ar, Xe, CO2, H2O and C6H6, and mercury porosimetry are characterized by much lower sensitivity (by several orders) than the QCM technique in respect to the studied weight per surface area unit. However, the adsorption isotherms of probe gases measured by standard methods allow one to investigate the structural characteristics of hard porous materials, including their porosity, pore size distribution, specific surface area, distribution of the adsorption potential and fractal dimension. Application of a combination of methods for the comprehensive characterization of surfaces, interfacial phenomena and adsorption of gases and solutes is fruitful. It should be noted that non-QCM gravimetric methods typically work in the gas phase and at reduced pressure, and this limitation is a barrier to investigations of biosystems in native liquid media. For many porous biomaterials used in tissue engineering which are soft hydrogels rather than rigid structures, these methods are hardly applicable, and other non-gravimetric methods are required.320

13.9 Dynamic contact angle measurements Contact angle between a sessile liquid droplet and a flat solid surface as a wetting phenomenon was originally described for non-porous and non-absorbing surfaces. Static contact angle  is defined geometrically as the angle formed by a liquid at the three-phase boundary where a liquid, gas and solid intersect. If the angle is less than 90ë the liquid is said to wet the solid. If it is greater than 90ë it is said to be non-wetting. A zero contact angle represents complete wetting.321±324 The static contact angle is measured directly by examining the shape of a sessile droplet on flat surface ± goniometry method. A sessile droplet on a tilted flat surface has two different angles ± advanced and receded. These angles are also obtained directly by examining geometry of the droplet. If the three-phase (liquid/solid/vapour) boundary is in actual motion, the angles produced are called dynamic contact angles (DCA) and are referred to as `advancing' and `receding' angles. They can be measured indirectly using microgravimetry ± tensiometry method. Tensiometry involves measuring the forces of interaction between a solid and a test liquid. The difference between the maximum (advanced/advancing) and minimum (receded/receding) contact angle values is called the contact angle hysteresis.325±328 A great deal of research has been dedicated to the analysis of the nature and significance of hysteresis. It has been used to characterize surface heterogeneity, roughness and mobility. Briefly, on non-homogeneous surfaces there will be domains which present barriers to the motion of the contact line. On chemically heterogeneous surfaces, these domains represent areas with different contact angles to the surrounding surface. For example when wetting with water occurs, hydrophobic domains will pin the motion of the

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contact line as the liquid advances thus increasing the contact angles. When water recedes, the hydrophilic domains will hold back the draining motion of the contact line thus decreasing the contact angles. From this analysis it can be seen that, when testing with water, advancing angles will be sensitive to hydrophobic domains and receding angles will characterize hydrophilic domains on the surface. For geometrically rough surfaces, the actual microscopic variations of the slope on the surface create the barriers, which pin the motion of the contact line and alter the macroscopic contact angles. Contact angle can also be considered in terms of the thermodynamics of the materials involved. This analysis involves the interfacial free energies between the three phases and is given by:

lv cos  ˆ sv ÿ sl

13:11

where lv , sv and sl refer to the interfacial energies of the liquid/vapour, solid/ vapour and solid/liquid interfaces. In the case of porous solids, powders and fabrics, the Washburn method is commonly used. If the forces of interaction, geometry of the solid and surface tension of the liquid are known, the contact angle may be calculated. The sample of the solid to be tested is hung on the balance and tared. The liquid is then raised to contact the solid. When the solid contacts the liquid, a sharp change of interaction forces is detected and a tensiometer registers this elevation as zero depth of immersion. As the solid is pushed into the liquid, Ftotal is measured by the microbalance: Ftotal ˆ wetting force ‡ weight of sample buoyancy force

13:12

Modern DCA equipment has software that tares the weight of the probe and removes the effects of the buoyancy force by extrapolating the graph back to zero depth of immersion. The remaining component is the wetting force defined as: Wetting force ˆ LV P cos 

13:13

where LV is the liquid surface tension, and P is the perimeter of the probe. This contact angle, which is obtained from data generated as the probe advances into the liquid, is the advancing contact angle. The sample is immersed to a set depth and the process is reversed. As the probe retreats from the liquid, the data collected is used to calculate the receding contact angle. The use of tensiometry for measurement of contact angle has several advantages over conventional goniometry. At any point on the immersion graph, all points along the perimeter of the solid at that depth contribute to the force measurement recorded. Thus, the force used to calculate  at any given depth of immersion is already an averaged value. One may calculate an averaged value for the entire length of the sample or average any part of the immersion graph data to assay changes in contact angle along the length of the sample. This technique allows one to analyze

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contact angles produced from wetting over an entire range of velocities from static to rapid wetting. The solid is brought into contact with the testing liquid and the mass of liquid absorbed into the solid is measured as a function of time. The amount absorbed is a function of the viscosity, density and surface tension of the liquid, the material constant of the solid, and the contact angle of the interaction. If the viscosity, density and surface tension of the liquid are known, the material constant and contact angle can be determined. The primary focus of contact angle studies is in assessing the wetting characteristics of solid/liquid interactions. The work of adhesion is defined as the work required to separate liquid and solid phases, or the negative free energy associated with the adhesion of solid and liquid phases. It is used to express the strength of the interaction between the two phases and is expressed by the Young-Dupre equation as: Wa ˆ …1 ‡ cos †

13:14

Work of cohesion is defined as the work required to separate a liquid into two parts, it is a measure of the strength of molecular interactions within the liquid. It is expressed as: Wc ˆ 2

13:15

Work of spreading is the negative free energy associated with spreading liquid over solid surface. Also referred to as a spreading coefficient it is given as: Ws ˆ …cos  ÿ 1†

13:16

Wetting tension as a measure of force/length is defined as: t ˆ Fw =P ˆ LV cos 

13:17

This value, wetting force normalized for length, also represents the product of the cosine of the contact angle and the surface tension. It allows for a characterization of the strength of the wetting interaction without separate measurement of surface tension, and is most helpful in characterizing multicomponent systems, where surface tension at an interface may not be equal to the equilibrium surface tension. Measurement of contact angles yield data, which reflect the thermodynamics of a liquid/solid interaction. Using a series of homologous liquids of differing surface tensions a graph of cos  vs is produced; it is found that the data form a line which approaches cos  ˆ 1 at a given value of . This is the maximum surface tension of a liquid that can completely wet the solid. This value, called the critical surface tension, is used to characterize a solid surface. Another way to characterize a solid surface is by calculating surface free energy, also referred to as the solid surface tension. This approach involves testing the solid against a series of well characterized wetting liquids. The liquids used must be characterized so that the polar and dispersive components of their surface tensions are known. The relevant equation derived by Owens and Wendt is:

Microgravimetry

l …1 ‡ cos †=… ld †1=2 ˆ … sp †1=2 ‰… lp †1=2 =… ld †1=2 Š ‡ … sd †1=2

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where l is the liquid surface tension and s is the solid surface tension, or free energy. The subscripts d and p refer to the dispersive and polar components of each parameter. The form of the equation is linear of the type y ˆ mx ‡ b. One can plot … lp †1=2 =… ld †1=2 vs l …1 ‡ cos †=… ld †1=2 . The slope gives … sp †1=2 and the y-intercept gives … sd †1=2 . The total free surface energy is merely the sum of its two component forces. In the case of liquid adsorption by a porous solid, a weight correction should be considered   x ÿ xZDOI ; 13:19 F ˆ p L cos q ‡ A
g…x ÿ XZDOI † ‡
m 1 ÿ xmax ÿ xZDOI where p is the object perimeter, L is the surface tension of a liquid, g is the acceleration due to gravity, x is the position of the plate and xZDOI is the position of the plate at zero depth of immersion,
m is the change in mass of a sample after complete withdrawal from the liquid. A method based on QCM for measuring sessile contact angles and surface energies of liquid-air and liquid-liquid interfaces has been described.329 This method involves measurement of the frequency change accompanying the introduction of a small liquid droplet to the center of a vibrating quartz resonator, which comprises an AT-cut quartz crystal sandwiched between two gold electrodes. If the density and viscosity of the liquid are known, the contact angle between the droplet and the gold substrate surface can be determined directly. QCM measurements of contact angles formed between aqueous droplets and gold surfaces modified with various organosulfur monolayers having different surface energies agree with sessile contact angles determined by optical goniometry. Furthermore, the QCM method can be used to measure the contact angle formed between an aqueous droplet and the QCM surface when both are submerged under an immiscible solvent such as hexadecane. In this case, the frequency change relies on the differences in the densities and viscosities of the water droplet and the fluid displaced by the droplet at the surface. The dependence of contact angle on the concentration of surfactant in a aqueous droplet provides for determination of the critical micelle concentration for aqueous phases in contact with air or an immiscible organic fluid. These measurements can be performed under conditions where contact angles cannot be measured readily, such as in the presence of opaque media or in the case of two liquids having similar refractive indexes.329 It is tempting to relate the surface energy of a solid to its biocompatibility, but our understanding of this relationship is poor.330±332 Even although sometimes it is possible to obtain a correlation between the surface energy of a biomaterial and its performance, as found for PMMA coated with a biomimetic phosphorylcholine coating,333,334 this relationship is far too complex to draw definitive conclusions.

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13.10 Conclusion Modern and most prospective developments of micro-, nano-, and femtogravimetry are based on the crystal sensor technology, which gives fruitful in-depth results for many complex systems, such as cells, biomacromolecules, polymers, thin films, drugs, and interfacial phenomena. Other gravimetric methods, direct (electronic balances, thermogravimetry and isothermal adsorption) or indirect (dynamic contact angle), are characterized by a much lower sensitivity than the crystal sensor techniques. There is an important trend in the applications of the crystal sensor techniques by combining with other methods such as AFM, SEM, TEM, SPR, XPS and specroscopy, that provides detailed information about complex interfacial phenomena, which will have both great theoretical and practical significance for the future.

13.11 Acknowledgements This work has been supported by DTI (UK) project MPP4.2. V. M. G.'s visit to the University of Brighton (UK) was supported by NATO (CLG Grant No. 979845).

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1256±9. 268. Muratsugu M, Kurosawa S and Kamo N, `Adsorption and desorption of F(ab')2 anti-human immunoglobulin G onto plasma polymerized allylamine film ± the application of the film to immunoassay', J Colloid Interf Sci, 1991 147(2) 378±86. 269. Aizawa H, Kurosawa S, Ogawa K, Yoshimoto M, Miyake J and Tanaka H, `Conventional diagnosis of C-reactive protein in serum using latex piezoelectric immunoassay', Sensor Actuat B ± Chem, 2001 76(1±3) 173±6. 270. Pavey K D, Ali Z, Olliff C J and Paul F, `Application of the quartz crystal microbalance to the monitoring of Staphylococcus epidermidis antigen-antibody agglutination', J Pharmaceut Biomed, 1999 20 (1±2) 241±5. 271. Zhou X C and Cao L, `High sensitivity microgravimetric biosensor for qualitative and quantitative diagnostic detection of polychlorinated dibenzo-p-dioxins', Analyst, 2001 126(1) 71±8. 272. Pùibyl J, Hepel M, HalaÂmek J and SklaÂdal P, `Development of piezoelectric immunosensors for competitive and direct determination of atrazine', Sensor Actuat B ± Chem, 2003) 91(1±3) 333±41. 273. J. Pùibyl, P. SklaÂdal, M. Hepel, Piezoelectric biosensors for polychlorinated biphenyls operating in aqueous and organic phases, QCM 2002, Acoustic Sensors Conference, Brighton, UK, 24±25 July 2002 274. Cooper M A, Dultsev F N, Minson T, Ostanin V P, Abell C and Klenerman D, `Direct and sensitive detection of a human virus by rupture event scanning', Nat Biotechnol, 2001 19(9) 833±7. 275. Liedberg B, Nylander C and LundstroÈm I, `Surface plasmon resonance for gas detection and biosensing', Sensor Actuat, 1983 4(2) 299±304. 276. Hutchinson A M, `Evanescent-wave biosensors ± real-time analysis of biomolecular interactions', Mol Biotechnol, 1995 3(1) 47±54. 277. Simpson T R E, Cook M J, Petty M C, Thorpe S C and Russell D A, `Surface plasmon resonance of self-assembled phthalocyanine monolayers: Possibilities for optical gas sensing', Analyst, 1996 121(10) 1501±5 278. O'Shannessy D J and Winzor D J, `Interpretation of deviations from pseudo-firstorder kinetic behavior in the characterization of ligand binding by biosensor technology', Anal Biochem, 1996 236(2) 275±83. 279. Holmes S D, May K, Johansson V, Markey F and Critchley L A, `Studies on the interaction of Staphylococcus aureus and Staphylococcus epidermidis with fibronectin using surface plasmon resonance (BIAcore)', J Microbiol Meth, 1997 28(1) 77±84. 280. Su X and Zhang J, `Comparison of surface plasmon resonance spectroscopy and quartz crystal microbalance for human IgE quantification', Sensor Actuat B ± Chem, 2004 100(3) 311±16. 281. Sellborn A, Andersson M, Fant C, Gretzer C and Elwing H, `Methods for research on immune complex activation on modified sensor surfaces', Colloid Surface B, 2003 27 295±301. 282. Hess C, Borgwarth K and Heinze J, `Integration of an electrochemical quartz crystal microbalance into a scanning electrochemical microscope for mechanistic studies of surface patterning reactions', Electrochim Acta, 2000 45(22±23) 3725±36. 283. Voinova MV, Rodahl M, Jonson M and Kasemo B, `Viscoelastic acoustic response of layered polymer films at fluid-solid interfaces: Continuum mechanics approach', Phys Scripta, 1999 59(5) 391±6.

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284. Narine S S, Hughes R and Slavin A J, `The use of inductively coupled plasma mass spectrometry to provide an absolute measurement of surface coverage, and comparison with the quartz crystal microbalance', Appl Surf Sci, 1999 137(1±4) 204±6. 285. Choi K H, Friedt J M, Frederix F, Campitelli A and Borghs G, `Simultaneous atomic force microscope and quartz crystal microbalance measurements: Investigation of human plasma fibrinogen adsorption', Appl Phys Lett, 2002 81(7) 1335±7. 286. Friedt J-M, Choi K H, Francis L and Campitelli A, `Simultaneous atomic force microscope and quartz crystal microbalance measurements: Interactions and displacement field of a quartz crystal microbalance', Jpn J Appl Phys Part I, 2002 41 (6A) 3974±7. 287. Smith A L and Shirazi H M, `Quartz microbalance microcalorimetry a new method for studying polymer-solvent thermodynamics', J Thermal Anal Calorim, 2000 59(1±2) 171±86. 288. Du J, Harding G L, Collings A F and Dencher P R, `An experimental study of Love-wave acoustic sensors operating in liquids', Sensor Actuat A ± Phys, 1997 60(1±3) 54±61. 289. Kalantar-Zadeh K, Trinchi A, Wlodarski W and Holland A, `A novel Love mode device based on a ZnO/ST-cut quartz crystal structure for sensing applications', Sensor Actuat A ± Phys, 2002 100(2±3) 135±43. 290. Cheeke J D N and Wang Z, `Acoustic wave gas sensors', Sensor Actuat B ± Chem, 1999 59(2±3) 146±53. 291. Zhang C, Caron J J and Vetelino J F, `The Bleustein-Gulyaev wave for liquid sensing applications', Sensor Actuat B ± Chem, 2001 76(1±3) 64±8. 292. Du J and Harding G L, `A multilayer structure for Love-mode acoustic sensors', Sensor Actuat A ± Phys, 1998 65(2±3) 152±9. 293. Andle J C, Vetelino J F, MW Lade M W and McAllister D J, `An acoustic plate mode biosensor', Sensor Actuat B ± Chem, 1992 8(2) 191±8. 294. Gizeli E, Stevenson A C, Goddard N J and Lowe C R, `A novel Love-plate acoustic sensor utilizing polymer overlayers', IEEE T Ultrason Ferr, 1992 39(5) 657±9. 295. Thompson D F and Auld B A, `Surface transverse wave propagation under metal strip gratings', Proc IEEE Ultrasonics Symp Williamsburg, VA, USA, 261±6, 1986. 296. DeÂjous C, Savart M, RebieÁre D and Pistre J, `A shear-horizontal acoustic plate mode (SH-APM) sensor for biological media', Sensor Actuat B ± Chem, 1995 27(1±3), 452±6. 297. Dahint R, Ros Seigel R, Harder P, Grunze M and Josse F, `Detection of non specific protein adsorption at artificial surfaces by the use of acoustic plate mode sensors', Sensor Actuat B ± Chem, 1996 36 (1±3), 497±505. 298. Gulyaev Y V, `Review of shear surface acoustic waves in solids', IEEE T Ultrason Ferr, 1998 45(4), 935±8. 299. Kovacs G, Vellekoop M J, Haueis R, Lubking G W and Venema A, `A Love sensor for (bio)chemical sensing in liquids', Sens Actuators A, 1994 43(1±3) 38±43. 300. Jakoby B and Vellekoop M J, `Analysis and optimisation of Love-wave liquid sensors', IEEE T Ultrason Ferr, 1998 45(5) 1293±302. 301. Du J, Harding G L, Ogilvy J A, Dencher P R and Lake M, `A study of Love-wave acoustic sensors', Sensor Actuat A ± Phys, 1996 56(3) 211±19. 302. Tamarin O, Comeau S, DeÂjous C, Moynet D, RebieÁre D, Bezian J and Pistre J,

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306. 307. 308. 309. 310. 311. 312. 313. 314. 315. 316. 317. 318. 319. 320. 321.

Surfaces and interfaces for biomaterials `Real time device for biosensing: design of a bacteriophage model using Love acoustic waves', Biosens Bioelectron, 2003 18(5±6) 755±63. Kalantar-Zadeh K, Wlodarski W, Chen Y Y, Fry B N and Galatsis K, `Novel Love mode surface acoustic wave based immunosensors', Sensor Actuat B ± Chem, 2003 91(1±3) 143±7. Shiokawa S and Kondoh J, `Surface acoustic wave sensors', Jpn J Appl Phys Part 1, 2004 43(5B) 2799±802. Joseph S, Gronewold T M A, Schlensog M D, Olbrich C, Quandt E, Famulok M, Schirner M, `Specific targeting of ultrasound contrast agent (USCA) for diagnostic application: an in vitro feasibility study based on SAW biosensor'. Biosens Bioelectron, (2004) doi: 10.1016/j.bios.2004.07.014, available online at www.sciencedirect.com Uttenthaler E, SchraÈml M, Mandel J and Drost S, `Ultrasensitive quartz crystal microbalance sensors for detection of M13-Phages in liquids', Biosens Bioelectron, 2001 16(9±12) 735±43. Wessa T, Barie N, Rapp M, Ache H J, `Polyimide, a new shielding layer for sensor applications', Sensor Actuat B ± Chem, 1998 53(1±2), 63±8. Barie N and Rapp M, `Covalent bound sensing layers on surface acoustic wave (SAW) biosensors', Biosens Bioelectron, 2001 16(9±12) 979±87. Wessa T, Rapp M and Ache H J, `New immobilization method for SAWbiosensors: covalent attachment of antibodies via CNBr', Biosens Bioelectron, 1999 14(1) 93±8. Harding G L, Du J, Dencher P R, Barnett D and Howe E, `Love wave acoustic immunosensor operating in liquid', Sensor Actuat A ± Phys, 1997 61(1±3) 279±86. O'Shannessy D J, Brigham-Burke M and Peck K, `Immobilization chemistries suitable for use in the BIAcore surface plasmon resonance detector', Anal Biochem, 1992 205(1) 132±6. Johnsson B, LoÈfaÊs S and Lindquist G, `Immobilization of proteins to a carboxymethyldextran-modified gold surface for biospecific interaction analysis in surface plasmon resonance sensors', Anal Biochem, 1991 198(2) 268±77. Galwey A K, `Is the science of thermal analysis kinetics based on solid foundations?: A literature appraisal', Thermochim Acta, 2004 413(1±2) 139±83. Carr N J and Galwey A K, `Decomposition reactions of solids (an experiment in reviewing)', Thermochim. Acta, (1984) 79 323±370. Galwey A K and Brown M E, Thermal Decomposition of Inorganic Solids, Amsterdam, Elsevier, 1999. Brown M E, Introduction to Thermal Analysis, 2nd edn, Dordrecht, The Netherlands, Kluwer Academic Publishers, 2001. Gallagher P K (ed.), Handbook of Thermal Analysis and Calorimetry, Amsterdam, Elsevier, 1998±2004. Adamson A W and Gast A P, Physical Chemistry of Surfaces, 6th edn, New York, Wiley, 1997. Gregg S J and Sing K S W, Adsorption, Surface Area and Porosity, 2nd edn, London, Academic Press, 1982. Tomlins P, Grant P, Mikhalovsky S, James S and Mikhalovska L, `Measurement of pore size and porosity of tissue scaffolds', Paper ID: J ASTM Int, 2004 1(1) JAI11510. Schrader M E and Loeb G, Modern Approach to Wettability, New York, Plenum

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Press, 1992. 322. Wu S, Polymer Interface & Adhesion, New York, Marcel Dekker, 1982. 323. Andrade J D, in Surface & Interfacial Aspects of Biomedical Polymers, Vol 1, New York, Plenum Press, 1985. 324. Bendure R L, `Dynamic adhesion tension measurement', J Colloid Interface Sci, 1973 42(1) 137±44. 325. Schwartz L W and Garoff S, `Contact-angle hysteresis on heterogeneous surfaces', Langmuir, 1985 1(2) 219±30. 326. Wang J H, Claesson P M, Parker J L and Yasuda H, `Dynamic contact angles and contact-angle hysteresis of plasma polymers', Langmuir, 1994 10(10) 3887±97. 327. Johnson R E, Dettre R H and Brandreth D A, `Dynamic contact angles and contact angle hysteresis', J Colloid Interface Sci, 1977 62(2) 205±12. 328. Drelich J, `Static contact angles for liquids at heterogeneous rigid solid surfaces', Polish J Chem, 1997 71(5) 525±49. 329. Lin Z X and Ward M D, `Determination of contact angles and surface tensions with the quartz crystal microbalance', Anal Chem, 1996 68(8) 1285±91. 330. Zisman W A, `Relation of the equilibrium contact angle to liquid and solid constitution', in: Contact Angle, Wettability, and Adhesion, Fowkes F M (ed.), Washington, DC, Amer Chem Soc, 1±51, 1963. 331. Ruckenstein E and Gourisankar S V, `A surface energeting criterion of blood compatibility of foreign surfaces', J Colloid Interface Sci, 1984 101(2) 436±51. 332. Schrader M E, `On adhesion of biological substances to low energy solid surfaces', J Colloid Interface Sci, 1982 88(1) 296±7. 333. Mikhalovsky S V, Santin M, Mikhalovska L I, Lloyd A W and Denyer S P, `Current trends in biomaterial coatings', in: Nanostructured Materials and Coatings for Biomedical and Sensor Applications, Gogotsi Y G and Uvarova I V (eds), Dodrecht, The Netherlands, Kluwer, 15±26, 2003. 334. Hayward J A and Chapman D, `Biomembrane surfaces as models for polymer design ± the potential for hemocompatibility', Biomaterials, 1984 5(3) 135±42.

Part III

Surface interaction and in-vitro studies

14

Interaction between biomaterials and cell tissues

Y I W A S A K I and N N A K A B A Y A S H I , Tokyo Medical and Dental University, Japan

14.1 Introduction In developing biomedical materials, we are concerned with their function, durability, and biocompatibility.1 Particularly, the surface properties of materials are directly related with this concern. Durability, particularly in a biological environment, is less well understood. Still, the tests we need to evaluate durability are clear. The important question in biocompatibility is how the device or material `transduces' its structural makeup to direct or influence the response of proteins, cells, and the organisms that are to relate to it. For devices and materials that do not leach undesirable substances in sufficient quantities to influence cells and tissues, this transduction occurs through the surface structure ± the body `recognizes' the surface structure and responds to it. For this reason we must understand the surfaces of biomaterials and the interfaces between biosubstances and biomaterials. In this chapter we describe the surface phenomena of biomaterials. First, the general recognition of a living organism to artificial materials and biocompatibility, which is a characteristic common to all materials, are discussed. Second, typical methods for surface analyses are introduced. Third, a design for `non-biofouling' surfaces, which are useful for materials that are exposed to blood or for diagnosis, is described. Finally, bioconnective materials based on the concept of a reliable system for connection to tooth substrates are discussed.

14.2 Surface properties of biomedical materials 14.2.1 Bioreactions on biomaterial surfaces2 A large amount of effort goes into the design, synthesis, and fabrication of biomaterials and related devices to ensure that they have appropriate mechanical properties, durability, and functionality. These functionalities of biomaterials and medical devices are primarily derived from the bulk structure of the material. In contrast, biological responses to materials are influenced by their

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14.1 Classification of interfacial biocompatibility.

surface chemistry and structure. Therefore, surface design and characterization are quite important in the development of biomaterials. Biocompatibility is a complementary definition for the required surface properties of biomaterials. `Biocompatibility is the ability of a material to perform with an appropriate host response in a specific application' (Williams, 1987).3 Thus, the meaning of biocompatibility depends on the specific application of the materials used. The roles of biomaterials can be divided into two categories, as shown in Fig. 14.1.4 One is `non-stimulative' and the other one is `bioconnective'. Because non-stimulative properties of materials complement non-activation, antithrombogenicity and tissue non-invasion are listed properties. These properties are necessary for materials that are exposed to blood. Conversely, bioconnective materials are required for adhesives for hard and soft tissue, sealants, and tissue engineering applications. Although these categorized materials will probably be used in completely different situations, the same surface properties, which can control interaction with the biocomponents, are required. Non-stimulative materials must reduce nonspecific protein adsorption and cell adhesion, socalled `biofouling.' Conversely, bioconnective materials should accumulate specific proteins and cells on their surfaces. However, it is usually difficult to control the interaction at the interface between artificial materials and a living organism. When these materials are exposed to a vital environment, non-specific protein adsorption and cell adhesion quickly occur. This biofouling reduces the functionality of the materials and also induces unexpected bioreactions such as thrombus formation, immune responses, complement activations, capsulations, etc.5,6 These reactions are important for maintaining vitality, however, they are serious problems for medical treatments that employ artificial materials. Most implanted medical devices serve their recipients well for extended periods by alleviating the conditions for which they were implanted, improving the quality of life and, with some device types, enhancing survival. However, some implants and extracorporeal devices ultimately develop complications that lead to device failure and thereby may cause harm to the patient or even death.

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Table 14.1 Typical bioreaction on artificial materials Protein adsorption Protein retention Lipid adsorption Bacteria adhesion Hemolysis Platelet adhesion Platelet activation Expression of new gene

Macrophage adhesion Phagocytosis Macrophage release Neutrophil activation Biodegradation Angiogenesis Cell spreading Fibrous encapsulation

All implants interact to some extent with the tissue environment in which they are placed. The biomaterial-tissue interactions encountered most frequently are summarized in Table 14.1. The important difficulty associated with medical devices is largely based on biomaterial-tissue interactions that include both the effects of the implant on the host tissue and the effects of the host on the implant. Several of the most important interactions in clinical and experimental implants and medical devices include inflammation and the `foreign body reaction,' immune response, systemic toxicity, thrombosis, device-related infection, and tumorigenesis. These interactions make up aberrations of physiological processes that function as common host defense mechanisms and are induced by non-specific biofouling. Thus, it is critical to prepare novel polymers that can control interactions with biocomponents at their interfaces.

14.2.2 Surface design of biomaterials1 To improve biointerfaces, the surface properties of biomaterials such as chemistry, wettability, topology, charge density, and bioactivity have been modified. Table 14.2 summarizes methods of surface modification and their efficiency. Coating is the most convenient method of surface modification. However, the affinity of a coating material to a substrate must be considered to reduce elution of the coating polymer. Graft polymerization is one of the familiar methods used to modify surface properties; the resulting modified surface is relatively stable because the graft polymers are connected with the substrate by covalent bonding. Graft polymerization can be achieved by several treatment methods such as application of chemical reagents, plasma-, corona-, photo-irradiation, etc. In particular, the corona discharge treatment is very convenient because it can be performed in an ambient atmosphere without the use of chemical reagents. Polymer blending is also an interesting method of surface modification in both research and technology. However, obtaining a homogeneous mixture of two or more polymers is difficult due to their high molecular weights. Moreover, the hydrophobic moiety shows a tendency to accumulate at the surface because

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Table 14.2 Surface modifications with polymer materials Modification methods Treatment

Effects

Chemical reaction

Oxygen functional groups Oxygen functional groups Sulfonic acid Thiol group Defluorine, oxygen functional groups Amination Oxygen functional groups

Chromic acid micture Chloric acid-sulfuric acid Sulfuric acid Sulfur Sodium Amine Sodium hypochlorite

Modifications of the original surface

Electron beam

Grafting

Radiation grafting Photografting Plasma, corona

Corona discharge Plasma etching

Depends (cross-linking, degradation) Oxygen functional groups Oxygen functional groups Various polymers

Noncovalent coatings Solvent coating Layer by layer adsorption Vapor deposition

Various polymers

Biological modification

Peptide, protein

Covalent bonding Intermolecular force

of the decrease in the surface free energy. In general, protein adsorption easily occurs on hydrophobic surfaces due to hydrophobic interaction. Although the stability of surface modification with blending would be better than that with coating, the disadvantage of blending is that surface properties are too difficult to control. An inter-penetrating polymer network (IPN) is defined as a network composed of two chemically independent cross-linked polymers. IPN technologies contribute properties from each of two independent polymers to a material. Monomers from the surface of a substrate penetrate and cross-link with the monomers in the subsurface of the substrate, giving a density-graded modification layer. This is a suitable method for surface modification because the reduction of the mechanical properties of the substrate can be reduced. There have been numerous studies for the fixing of bioactive molecules to materials to give their surfaces biological functions. For non-stimulative materials, particularly, blood antithrombogenic surfaces, heparin and thrombomoduline were fixed. In contrast, cell adhesive peptides and proteins were also immobilized on conventional polymer surfaces to improve affinity to cells. As expected, some of these bioactive molecules worked on the surface for short periods. However, their activity often declined meaning that they are inadequate for use in long-term medical treatment. In addition, to obtain bioactive molecules with high efficiency, the surface designs of the substrates

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are important because nonspecific biofouling and interaction with the substrate also reduce functionality.

14.3 Surface analyses of biomedical materials1 14.3.1 Introduction Surface analysis is quite important for understanding the properties of biomaterials. Of course, the methods of surface analyses of materials are constantly being improved. Table 14.3 shows the characteristics of many common methods of surface analysis, including the depth of analysis and the spatial resolution. For surface analysis, we must consider several points. The surface structure of the material is often mobile. The movement of the atoms and molecules near the surface in response to the outside environment is often highly significant. In response to a hydrophobic environment, more hydrophobic components may migrate to the surface of a material. In some methods of surface analysis, materials are kept under a condition of an ultra high vacuum. This condition is completely different from a physiological environment. Characterization of a Table 14.3 Typical surface analysis methods for biomaterials Method

Principle

Depth analyzed

Area analyzed

Contact angles

Liquid wetting of surfaces is used to estimate the surface energy

0.3±2 nm

1 mm

XPS: X-ray photoX-rays cause the emission of electron spectroscopy electrons of characteristic energy

1±25 nm 10±150 m

SIMS: secondary ion mass spectroscopy

Ion bombardment leads to the 1 nm±1 m emission of surface secondary ion

10 nm

FTIR-ATR: Fourier transform infra-red spectroscopy, attenuated total reflectance

IR radiation is adsorbed in exciting molecular vibrations

1±5 m

10 m

SPM: scanning probe Measurement of the quantum microscopy tunneling current (STM) or van der Waals repulsion (AFM) between tip and surface

0.5 nm

0.1 nm

SEM: scanning electron microscopy

0.5 nm

4 nm

0.3 m

2 m

Secondary electron emission caused by focused electron beam is measured and spatially imaged

SPR: surface plasmon Measurement change in the resonance refractive index in evanescent field generated surface

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Surfaces and interfaces for biomaterials

material surface under such conditions must be made very carefully because the actual surface has hydrophilic and hydrophobic components. Surfaces readily contaminate. Under ultra-high vacuum conditions, this contamination can be retarded. However, in view of the atmospheric pressure conditions under which all biomedical devices are used, we must learn to live with some degree of contamination. The significant questions here are whether we can make devices that have a constant, controlled level of contamination and can avoid undesirable contamination. This is critical in order that a laboratory experiment on a biomaterial generates the same results when repeated after one day, one week, or one year, and that a biomedical device performs for the physician in a constant manner over a reasonable operating life. Some analytical methods also apply to the characterization of a biological surface that forms an interface with an artificial material. Compared with an artificial surface, a biological surface is more complex, flexible, and easier to contaminate. To characterize the biological surface, we must then be cognizant of the circumstances under which the biointerface is used. A wide variety of methods has therefore to be applied for understanding the reliable surface properties of materials. Here, a few typical methods for the surface analysis of biomaterials are introduced.

14.3.2 Contact angle method The force balance between the liquid-vapor surface tension ( LV ) of a drop of liquid and the interfacial tension between a solid and the drop ( SL ), which is manifested through the contact angle () of the drop with the surface (Fig. 14.2), can be used to characterize the energy of the surface ( SV ). The basic relationship describing this force balance is:

SV ˆ SL ‡ LV
cos 

14:1

The energy of the surface, which is directly related to its wettability, is a useful parameter that has often correlated strongly with biological interaction. Unfortunately, SV cannot be obtained directly since this equation contains two unknowns, SL and SV . Therefore, SV is usually approximated by the

14.2 Water contact angle measurement using a sessile drop technique.

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14.3 Schematic representation of corona irradiation apparatus.

Zisman method for obtaining the critical surface tension.7 Experimentally, there are several ways to measure the contact angle. From this measurement, a fundamental characterization of material surface can be obtained. The corona discharge treatment is one of the most robust methods for surface modification of polymer materials to control surface wettability and produce graft polymers. Lee and co-workers prepared a wettability gradient surface on conventional polymer materials and investigated protein or cell/surface interactions.8 Figure 14.3 is a schematic diagram showing the corona discharge apparatus for the preparation of a gradient polyethylene sheet. Peroxide is produced by the corona discharge treatment, and this peroxide transfers to various functional groups with oxygen. Figure 14.4 shows the water contact angles of the PE surface after treatment with corona irradiation by changing the energy along the sample distance. The water contact angle is decreased with an increase in the energy of the corona irradiation.

14.4 Change in water contact angle of PE surface by corona treatment.

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Surfaces and interfaces for biomaterials

14.3.3 X-ray photoelectron spectroscopy (XPS)9 X-ray photoelectron spectroscopy (XPS) is generally regarded as an important and key technique for the surface characterization and analysis of biomedical polymers. This technique, also called ESCA, provides a total elemental analysis, Ê of any solid surface that is except for hydrogen and helium, of the top 10±200 A vacuum stable or can be made vacuum stable by cooling. Chemical bonding information is also provided. Of all the presently available instrumental techniques for surface analysis, XPS is generally regarded as being the most quantitative, the most readily interpretable, and the most informative with regard to chemical information. For these reasons, it has been highly recommended and used by biomedical researchers for the analysis of medical materials. The basic principle of XPS is the photoelectric effect, the phenomenon for which Einstein received the Nobel Prize. Three possibilities are evident. 1. 2. 3.

The photon may traverse through the atom without significantly interacting with either the orbital electrons or the nucleus. The photon may be scattered by the atomic orbital electron resulting in a partial loss of photon energy. This process is called Compton scattering. The photon may interact with the atomic orbital electron such that there is total and complete transfer of the photon's energy to the electron. This is the basic process in XPS.

Given that the photon energy is greater than the binding energy of the electron in the atom, the electron is then ejected from the atom with a kinetic energy approximately equal to the difference between the photon energy and the binding energy (Fig. 14.5). Therefore, the basic equation for XPS is: Eb ˆ h ÿ Ek

14:2

where Eb is the electron binding energy, Ek is the electron kinetic energy measured by the instrument, and h is the photon energy (h is Planck's constant and  is the X-ray frequency). All energies are usually expressed in electron volts (eV). Measuring the kinetic energy enables calculation of the binding energy. By knowing the binding energy, we can identify the atom.

14.5 Schematic views of XPS analysis and interaction of X-ray photon of energy h with an atomic orbital electron.

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14.6 XPS spectra of the gradient polymer surfaces at positions 0.5 cm, 2.5 cm, and 4.5 cm from the starting point of corona treatment.

The peroxides produced on a polymer substrate by corona irradiation (Fig. 14.3) work as an initiator for free-radical polymerization. When the coronairradiated polymer film is soaked in a monomer solution, the graft polymers were obtained from the film surface. Figure 14.6 shows the XPS spectra of C1s and P2p on a gradient poly(2-methacryloyloxyethyl phosphorylcholine) (PMPC) grafted polyethylene surface.10 The concentration of the MPC unit on the surface increased with an increase in the distance from the starting point of the corona irradiation. That is, a phosphorus peak attributed to the phosphorylcholine group in the MPC unit was found at 134.5. Moreover, a carbonyl peak was seen at 288.5 eV because it was clear compared with the untreated PE surface. The protein adsorption and cell adhesion decreased with an increase in the density of the grafted poly(MPC).

14.3.4 Scanning probe microscopy (SPM)11 A number of methods that provide information about the structure of a solid surface, its composition, and current oxidation states have come into use. The recent upsurge of activity in scanning probe microscopy has resulted in

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Surfaces and interfaces for biomaterials

14.7 Schematic diagram of the scanning tunneling microscope (STM) and atomic force microscope (AFM).

investigation of a wide variety of surface structures under a range of conditions. The ability to control the position of a fine tip in order to scan surfaces with subatomic resolution has brought scanning probe microscopy to the forefront in surface imaging techniques. There are two primary techniques; scanning tunnel microscopy (STM) and atomic force microscopy (AFM), as illustrated in Fig. 14.7. STM relies on measurement of the experimentally decaying tunneling current between a metal tip and the conducting substrate. Since its development in the early 1980s and the recognition of its inventors with the presentation of the 1986 Nobel Prize, STM has found wide use in studies of both inorganic and organic materials. Supporting a fine tip, a piezoelectrically positioned cantilever spring provides the means of measuring surface forces in the range of 10ÿ13 to 10ÿ6 N. The AFM measures deflections in the cantilever due to capillary, electrostatic, van der Waals, and frictional forces between the tip and the surface. Not limited to conductive surfaces, AFM measurements can also be made on organic and inorganic surfaces and on surfaces immersed in liquids. Conical tips of silicon have points of 5 to 50 nm (radius of curvature). However, numerous probes have been used including attaching a colloidal particle of several micrometers to a cantilever, as described by Butt and co-worker12 and Ducker and co-worker.13 Since AFM measures force, it can be used with both conductive and nonconductive specimens. Since force must be applied to bend a lever, AFM is subject to artifacts caused by damage to fragile structures on the surface. Both STM and AFM methods can function well for specimens under water, in air, or under a vacuum. For exploring biomolecules or mobile organic surfaces, the `pushing around' of structures by the tip is a significant concern. With both methods, it is difficult to achieve good-quality, reproducible images of organic substrates. However, some of the successes to date are exciting enough that the future of these methods in biomedical research is ensured. Figure 14.8 is an AFM image of a DPPC liposome adsorbed on a 2methacryloyloxyethyl phosphorylcholine (MPC) polymer surface. Using AFM, we can observe biomolecules and their association with the native structure.14

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14.8 AFM image of liposomes adsorbed on the MPC polymer.

14.4 Design for non-biofouling surface An assortment of materials have been used for manufacturing medical devices including artificial organs that are exposed to blood. However, the only polymers presently used for this type of application are conventional materials such as poly(vinyl chloride) (PVC), polyethylene(PE), poly[methyl methacrylate (MMA)], segmented polyether urethane (SPU), poly(dimethylsiloxane), poly(tetrafluoroethylene), cellulose, and polysulfone (PSF). Because nonspecific protein adsorption occurs on these materials, they cannot reduce the level of reactions to foreign bodies. Therefore, drug infusion is required during clinical treatments using these medical devices to prevent reactions to foreign bodies. To reduce the level of bioreaction on a surface, polymeric surfaces, which can reduce nonspecific protein adsorption, that is, `nonfouling surfaces' have been studied. The molecular design of nonfouling polymers for biomedical

Poly(2-hydroxyethyl methacrylate) Poly(acrylamide) Block-type copolymer Graft-type copolymer Segmented polyurethane Poly(ethylene glycol)

Hydrophilic surface zero interfacial free energy

Heterogenic surface microdomain concept

Molecular cilia mechanism

Negative charged surface Sulfonated polymer - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - Synthetic polymer Pseudomembrane formation Expanded PTFE + Poly(ethylene terephthalate) Biologically active Immobilization of bioactive protein polymer Heparinization Heparine releasing polymer Heparinized polymer Urokinase Thrombomoduline - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - - Synthetic polymer Albumin adsorbed surface Polyurethane with alkyl group + Biological Phospholipid adsorbed surface Phospholipid polymer molecules biomembrane-like surface formation

Polytetrafluoroethylene(PTFE) Polydimethylsiloxane Polyethylene

Hydrophobic surface zero critical surface theory

Synthetic polymer

Typical polymer

Basic concept

Type of materials

Table 14.4 Surface immobilizations for blood contacting materials

s

Biological approach

Biochemical approach

Interfacial chemical approach

Physicochemical approach

s s

s

Approach

ss

s

s

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14.9 Chemical structure of MPC.

applications is classified into the four categories listed in Table 14.4. One of the most effective methods for producing a nonfouling surface is the modification of conventional materials with polymers having a phospholipid polar group mimicking a biomembrane surface. In 1978, Nakabayashi designed a methacrylate monomer with a phospholipid polar group, 2-methacryloyloxyethyl phosphorylcholine (MPC) to obtain new medical polymer materials (Fig. 14.9).15 However, at that time, the purity and yield of the MPC was not sufficient to evaluate their functions. Ishihara improved the synthetic route of the MPC16 and prepared a wide variety of polymers containing the MPC unit.17±19 The homopolymer of the MPC is soluble in water and the solubility of the MPC polymers can easily be controlled by changing the structure and fraction of the comonomer. Many papers have described the effect of the MPC moiety on the nonfouling property using poly(MPC-co-n-butyl methacrylate(BMA)) (PMB). Figure 14.10 shows experimental columns containing polymer-coated poly(methyl methacrylate) beads and micrographs of the bead surface after contact with human blood without an anticoagulant. On the hydrophobic poly[nbutyl methacrylate (BMA)], many adherent cells can be observed. Moreover, the activation and aggregation of the cells and clot formation can be observed on the

14.10 Blood compatibility test on PBMA and PMB30 with microsphere column method.

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14.11 Amount of protein adsorbed on polymer surfaces from human plasma.

poly(BMA). In contrast, poly(MPC-co-BMA) (PMB) with a 30 mol% MPC unit (PMB30) effectively suppressed cell adhesion. The blood compatibility of MPC polymer is due to the surface property that reduces nonspecific protein adsorption. The protein adsorption-resistance on MPC polymer surface has been studied with consideration of not only the amount of adsorbed proteins but also of the species of the proteins. Figure 14.11 shows the amount of protein adsorbed on the PMB, poly(HEMA), and poly(BMA) after contact with plasma for 60 min.20 On the poly(BMA), many more proteins were adsorbed compared with those on the poly(HEMA) and the PMB. The amount of proteins adsorbed on the PMB decreased with an increase in the amount of MPC in the polymer. The species and distribution of the proteins adsorbed on the PMB were also determined by gold colloid- and radio-labeled immunoassays. 21 These experiments demonstrated that the PMB could reduce plasma protein adsorption nonspecifically. The thrombus formation on the conventional polymeric materials occurred through the multilayers of plasma proteins denaturated by contact with the surfaces. The secondary structure of bovine serum albumin (BSA) and bovine plasma fibrinogen (BPF) adsorbed on the PMB was evaluated by circular dichroism (CD) spectroscopy.22,23 Figure 14.12 shows the CD spectra of BSA in PBS and that adsorbed on the polymer surface. For the BSA in PBS, the mean molecular residual ellipticity, [], was a large negative value at 222 nm. The CD spectrum of BSA adsorbed on

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14.12 CD spectra of BSA in PBS and that on adsorbed polymer surfaces. (A): poly(HEMA), (B): PMB10, (C): PMB30, (D): BSA/PBS.

PMB was almost the same as that in PBS. The negative ellipticity at 222 nm of BSA adsorbed on the MPC polymers increased with a decrease in the ratio of MPC, then neared zero in the case of BSA adsorbed on poly(HEMA). We were able to find the same tendency for BPF. Calculation of the -helix content of BSA and BPF revealed that the PMB could effectively suppress the conformational change of the proteins even when the proteins were adsorbed on the surface. Conversely, it could be seen that the -helix content of both proteins adsorbed on poly(HEMA) decreased significantly. It is thought that the MPC polymer suppresses protein adsorption. Recently, it was reported that the structure of the water absorbed in/on the polymer materials influences protein adsorption on their surfaces. The water structure in hydrated polymers was determined to enhance the resistant properties of protein adsorption on surfaces containing phospholipids. Park et al. proposed a very important model for protein adsorption on polymer surfaces. The adsorption of proteins on a polymer surface via hydrophobic interaction requires an exchange of bound water between the protein and the surface.24 Therefore, the amount of bound water might be the important parameter in understanding protein adsorption. Tsuruta reported that the random networks of water molecules on material surfaces were very important in explaining the protein adsorption that occurs on them.25 Protein adsorption processes are considered to start with protein trapping by random networks of water molecules on the material surface. The material surface, which cannot construct hydrogen bonding with water, will then reduce protein adsorption. Table 14.5 lists the free water concentration in a hydrated

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Surfaces and interfaces for biomaterials Table 14.5 Characteristics of hydration state of polymer Poly(HEMA) Heqa Free water fraction at Heq at H=0.36 a

Heq ˆ

0.40 0.34 0.28

10

PMB

0.23 0.25

30 0.84 0.84 0.69

Weight of water in the polymer membrane at 25 ëC Weight of polymer membrane saturated with water

polymer membrane with a 0.36 water fraction determined by differential scanning calorimetry.23 The fraction of free water (not bound water) in the MPC polymer was 0.65, which was found to be significantly higher than that in poly(HEMA), which was 0.28. Furthermore, the structure and hydrogen bonding of water near the PMB were analyzed in their aqueous solutions and in thin films with contours of O-H stretching of Raman and attenuated total reflection infrared (ATR-IR) spectra, respectively.26,27 The relative intensity of the collective band (C value) corresponding to a long-range coupling of O-H stretchings of the Raman spectra for the aqueous solution of PMB was very close to that for pure water, which is in contrast with the smaller C value in the aqueous solution of ordinary polyelectrolytes. A similar tendency was also observed on hydrated thin polymer films. These results suggest that PMB does not significantly disturb hydrogen bonding between water molecules in either the aqueous solution or the thin film systems. The equilibrium amount of proteins, bovine serum albumin (BSA), and bovine plasma fibrinogen (BPF) adsorbed on the polymer surface was measured and represented with a free water fraction in the hydrated polymers, as shown in Fig. 14.13.28 The amount of both proteins adsorbed on poly(HEMA), poly(acryl amide (AAm)-co-BMA), and poly(N-vinylpyrrolidone(VPy)-co-BMA) were larger than those on the poly(MPC-co-dodecyl methacrylate)(PMD) and PMB. It was reported that the theoretical amount of BSA and BPF adsorbed on the surface in a monolayer state are 0.9 and 1.7 g/cm2, respectively. On the surfaces of MPC polymers, the amount of adsorbed proteins was less than these theoretical values. A schematic representation of the nonfouling property of PMB surfaces is shown in Fig. 14.14. When the PMB comes in contact with biofluids containing proteins, the hydrated surface of the PMB has a high free-water content and cannot be recognized by proteins. Cell adhesion is also reduced on these surfaces because there is no adhesive protein on the surfaces. Recently, it was found that the MPC polymer surface could reduce not only biofouling but also the inflammatory responses of cells in contact with the surface. Therefore, compared with other nonfouling surfaces, the MPC polymer surface might offer many advantages for biomedical materials.

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14.13 Relationship between free water fraction in hydrated polymer membrane and amount of proteins adsorbed on the polymer. :BSA, [BSA] in PBS ˆ 0.45 g/dl, :BPF, [BPF] ˆ 0.30 g/dl.

14.14 Possible mechanism of nonfouling property on MPC polymer surface.

MPC is a robust compound for surface modification to improve biocompatibility because it can be used in a wide variety of technologies. Figure 14.15 shows the surface modification of conventional polymer materials with MPC. Blood-compatible biomedical devices employing these technologies have been explored.

14.5 How to connect tissues with biomaterials There is a great demand for connecting tissues to devices, for artificial organs, and for biomaterials to support organ functions and injured and/or unwell tissues. It is not so difficult to unite natural tissues, which could heal by

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14.15 Various MPC polymer immobilizations on conventional polymer surface.

themselves. However, this is not always possible and we have to assist organ and tissue functions with biomaterials. The ability to introduce blood to a dialyzer three times a week is required for patients with chronic kidney ailments to keep them alive. Artificial blood vessels must be sutured to natural blood vessels because there is no suitable adhesive to connect biomaterials and tissues. Infection at the interface is a severe problem in medical treatment. Living tissues generally reject biomaterials in an attempt to avoid adverse effects. Nevertheless, we are attempting to apply biomaterials to remedy an organ failure such as the use of an artificial kidney. Adequate blood access is a critical problem for patients with chronic kidney disease. In several senses, biocompatibility is required at the tissue/biomaterial interface. When we consider the mechanism of pannus formation at the natural vessel side of connected artificial and natural blood vessels working for a long time, the differences in their biomechanical properties must be considered. The pulsing of blood vessels causes severe stress when the mechanical properties are different. The natural vessels change to minimize the adverse effect of the artificial vessels and their thickness increases to accommodate the situation. `Adhesive' is a general term that covers designations such as cement, glue, paste, fixative, and bonding agents used in the many areas of adhesive technology. Adhesive systems consist of one- or two-part organic and/or inorganic formulations that set or harden through the action of several different mechanisms. The applications of adhesive biomaterials range from soft tissue adhesives used both externally to temporarily attach accessory devices such as colostomy bags and internally for wound closure and sealing, to hard tissue adhesives used to bond prosthetic materials to teeth and bone on a more permanent basis. All of these physiological environments are hostile, and a major problem in the formulation of medical and dental adhesives is to develop

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a material that will be easy to manipulate, interact intimately with the tissue to form a suitable bond, and be biocompatible. Over the past two decades, more success at a clinical level has been achieved in bonding hard tissues than soft tissues. The tooth surface is an important surface, which should be given special note. Dental biomaterials are applied to rehabilitate decayed areas formed on dental hard tissues, because these tissues do not heal and regenerate. Completion of wound healing is essential to inhibit infection at the surface of the injured tissue before applying any biomaterials. Unfortunately, teeth do not heal by themselves because there are no blood vessels to initiate the wound-healing process. Dentists have been affixing restorations to prepared cavities and abutments as ordinary dental treatment for a long time. However, the treatment often does not survive on the tissues and the prostheses soon detach. It has been hypothesized that a good adhesive for dental hard tissues could provide good resolution in improving dental treatment by increasing the retentive force combined with the adhesive force. However, connecting natural tissues, even dental hard tissues, to biomaterials is not so easy. Dentists believe that the detachment of a restoration occurs by `microleakage' and `secondary caries' but they could not explain these mechanisms exactly. They misunderstood that dental cements cured by acid and base reactions are soluble in saliva. Therefore, they tried to minimize the dissolution of cured zinc phosphate cement luted on the hard tissues and accurately fabricated restorations with the prepared tissues by decreasing the thickness of the cement. They developed zinc carboxylate cements and glass-ionomer cements to decrease the solubility in saliva. They should understand that the cements dissolved if the abutments were dentin but not if the abutments were enamel. There are big differences between the two types of tissue. Cured cements can intimately attach to enamel etched with acidic cement pastes before they cure. In contrast, they are attached to dentin demineralized with acidic cement pastes before they have cured. The demineralized dentin is permeable to saliva, which dissolves the cement. In 1955, Buonocore developed a method of bonding self-cured acrylic resin to etched enamel. This technology was introduced in orthodontic treatment to affix plastic and metal brackets directly to provide stress on teeth in order to align them. Bonding to enamel is believed to be not as difficult as bonding to dentin, and the etching of the enamel gives a good bonding surface. Enamel does not contain as much water and peptides as dentin does; the principal constituents are hydroxyapatite crystals. Tag formation on the etched enamel surface gives good mechanical retention, which provides bond strength to the substrate. Therefore, stronger etching agents are used to make bigger tags to create stronger bonds to enamel. However, aggressive etchants weaken enamel. The bonding mechanism is explained in the illustration in Fig. 14.16.29 During studies to improve the direct bonding system in orthodontic treatment, it

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14.16 Schematic diagram of hybrid concept for resin bonding to enamel.

was found that the length of the tags changes with the adhesive used. Adhesives containing a hydrophobic and hydrophilic methacrylate appeared to penetrate beyond the etched surface and encapsulate the prisms. The hybrid zone does not increase the bonding strength significantly, but does promote stability due to a supplemental thin layer of resin-reinforced tissue. This helps minimize the adverse effect of etching the enamel and gives the enamel acid resistance. Several mechanisms have been proposed for bonding to dentin. Fixation of restorative materials had been performed by curing fluid substances on the surface of the tooth, which is on the outside of the tooth. Many dental researchers thought adhesive bonding was the result of surface phenomena. Consequently, wettability and chemical reaction were thought to be important in creating bonds to dentin. Unfortunately, these considerations have not worked in developing reliable systems for bonding to dentin. In 1982, Nakabayashi proposed a new mechanism to enable the bonding of resin to dentin;30 this system was widely accepted in the 1990s. The system is described as being micromechanical at the molecular level. Monomers having both hydrophobic and hydrophilic groups are impregnated by the exposed collagen of demineralized superficial dentin, which has a good capacity to accept the diffusion of monomers. Their polymerization in situ forms copolymers entangled with the collagen fibers and encapsulated hydroxyapatite crystalline. The entanglement locks the dentin with cured resin, which affords

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the bonding strength. It is the generation of a `hybridized dentin,' which is a transitional zone of resin-reinforced dentin, sandwiched between cured resin and the unaltered dentin substrate, and a mixture of dentinal components and cured resin at the molecular level. It is a kind of functional graded material. The hybrid layer is resin-infiltrated enamel, dentin, or cementum. The chemical and physical properties of these zones are quite different from those of the original tooth structure because the tissue has been partially demineralized and then impregnated with resin. The resulting structure is neither resin nor tooth but a hybrid of the two. It is not located on the surface but prepared within the subsurface of the substrates. The zones are so-called tissue-engineered demineralized dentin with polymers. The mineral phase of the hard tissue is purposely dissolved to create a diffusion pathway to the monomers. This matrix is then infiltrated with monomers to intentionally change their physical and chemical properties. The clinical success of pit and fissure sealants on enamel indicates how acid resistant these resin-enamel hybridizations can be over many years. The infiltrating resin actually envelops apatitic crystallites in enamel to improve their acid resistance. This is made possible by the pretreatment of enamel with acids, which increases the surface roughness on a microscopic scale. Indeed, resinbonded enamel contains micro-tags of resin prepared between the crystalline and enamel rods etched at a diameter as small as the molecular level, as illustrated in Fig. 14.16. It is essential to know the importance of hybridized dentin in dental treatment more than the bonding of restorations to tooth structures. We have believed that caries are formed by infections of acid-producing microorganisms living in our mouths. This is true in a sense, but we must note that these microorganisms do not infect dental tissues from the beginning of tooth development. Acids, mostly lactic acid and their related products, react with basic calcium phosphate, hydoxyapatite, and dissolve the tissues. It is a simple neutralization and demineralization of enamel and dentin. The demineralized enamel surface, without visible caries, is restored by the recrystallization of the hydroxyapatite present in the saliva. If the rate of demineralization does not exceed that of recrystallization, the enamel is sufficiently stable dynamically. However, if the rate of the former is faster than that of the latter, caries should develop. Consequently, it is very important to clean our teeth after eating food, which helps microorganisms form a plaque biofilm. The development of caries is evidence of a good acid delivery system to the enamel. Enamel is a very important tissue that protects the health of our teeth by acting as a barrier to protect the dentin and underlying pulp. It does this by inhibiting the invasion of lactic acid to the dentin to form dentin caries. Once dentin is exposed, it is impossible for dentists to heal it using dental materials, which has been a basic misunderstanding. The pulp in exposed dentin must accept the invasion of stimuli and microorganisms, which irritate the pulp. To

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prevent septicemia, sometimes we have to kill the tooth if the pulp must be extracted. Non-vital teeth constitute dead tissue. We have to prevent it by applying artificial enamel at the early stages of caries development and inhibit the invasion of lactic acid as our teeth cannot inhibit it. The healing of wounds of exposed dentin caused by caries is impossible and we must naturally extract the tooth. Although, dentists have tried to heal caries, this is impossible, as mentioned before. Blood coagulation after bleeding could initiate gingival wound healing and protect us from further infection. As is enamel, hybridized dentin is an impermeable barrier to lactic acid and can inhibit further demineralization of exposed dentin. We have been suffering secondary caries and detachment of prostheses affixed to abutments. Dental researchers have worked hard to inhibit microleakage by several methods but resolution was difficult. The meaning of microleakage, which damages restored teeth, is the marginal leakage of lactic acid, the principal product of acidproducing microorganisms in the mouth, and demineralization of abutments by the acid. We could say that hybridized dentin is artificial enamel that can protect dentin from further demineralization caused by acid. The preparation of hybridized dentin could be a man-made process of wound healing, pseudowound healing, of exposed dentin because it can protect further demineralization leaving only healed tissue. We could then easily provide rehabilitation by affixing prostheses to reconstruct the esthetics and functions of lost tissues. Dentists could attach the restoration to the top of the hybridized dentin because the top surface has good affinity for attachment with resin composites and resin cements. The interaction of tissues and biomaterials at biologic interfaces is extremely important but it is very difficult to connect natural tissues with artificial materials. The bonding of resins to mineralized tissues provides a good example of such an interface. The three major hard tissues are enamel, dentin, and bone. Unlike bone, the former two, the hard tissues of the teeth, do not regenerate. When biocompatible materials, such as dental implants, are placed close to bone, hydroxyapatite prepared by the surrounding bone fills the intervening space and a connection is formed. In contrast, when such a material is placed close to a tooth, no such connection develops because there is no biologically active soft tissue between the two structures to promote regeneration (Fig. 14.17). Unfortunately, hybridization of bone with impregnated adhesives killed bone-inducing cells by their encapsulation, and the bonding mechanism could not be applied to develop an adhesive bone cement between bone and prostheses.

14.6 Conclusion Surface design and characterization as well as mechanical properties are quite important in the preparation of biomaterials. The moment a living body is

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14.17 Interactions of biocompatible materials with tooth, bone, and soft tissue.

exposed to a material surface, it determines whether the material can be accepted. Requirements of biomaterial surfaces vary depending on their use and the environment in which they are used. However, biocompatibility is a necessary element for every biomaterial. Concerning biological structure and function, and applying various analyses for surface characterization may contribute to the exploration of suitable material surfaces.

14.7 References 1. Ratner B D (1996), 'Surface properties of materials,' in Ratner B D, Hoffman A S, Schoen F J, and Lemons J E, Biomaterials Science, San Diego, Academic Press, 21± 34. 2. Schoen F J and Anderson J M (1996), `Host reactions to biomaterials and their evaluation,' in Ratner B D, Hoffman A S, Schoen F J, and Lemons J E, Biomaterials Science, San Diego, Academic Press, 165±173. 3. Williams D F (1987), `Definitions in Biomaterials'. Proceedings of a Consensus of the European Society for Biomaterials, Chester, England, March 3±5 1986, New York, Elsevier. 4. Lee J W and Gardella J A Jr. (2002), `Surface perspectives in the biomedical applications of poly(alpha-hydroxy acid)s and their associated copolymers,' Anal Bioanal Chem, 373, 526±573. 5. Brash J L and Horbett T A (1987), Proteins at Interfaces ± Physicochemical and Biochemical Studies, ACS Symposium Series 343, Washington, DC, American Chemical Society. 6. Brash J L and Horbett T A (1995), Proteins at Interfaces II ± Fundamentals and Applications, ACS Symposium Series 602, Washington, DC, American Chemical Society. 7. Zisman W A (1964), `Relation of the equilibrium contact angle to liquid and solid constitution,' in Fowkes F M, Contact angle, wettability and adhesion, ACS Advances in Chemistry Series 43, Washington, DC, American Chemical Society, 1±51.

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8. Lee J H and Lee H B (1993), `A wettability gradient as a tool to study protein adsorption and cell adhesion on polymer surfaces,' J Biomater Sci Polym Ed, 4, 467±81. 9. Andrade J D (1985), Surface and Interfacial Aspects of Biomedical Polymers, vol. 1, Surface chemistry and Physics, New York, Plenum Press. 10. Iwasaki Y, Ishihara K, Nakabayashi N, Khang G, Jeon J H, Lee J W, and Lee H B (1998), `Platelet adhesion on the gradient surfaces grafted with phospholipid polymer,' J Biomater Sci Polym Ed, 9, 801±816. 11. Adamson A W and Gast A P (1997), Physical chemistry of surfaces, New York, John Wiley & Sons, Inc. 12. Butt H J, Jaschke M, and Ducker W (1995), `Measuring surface forces in aqueous electrolyte solution with the atomic force microscope,' Bioelectrochem Bioenerg, 38, 191±201. 13. Ducker W A, Xu Z (1994), Israelachvili J N, `Measurement of Hydrophobic and DLVO Forces in Bubble-Surface Interactions in Aqueous Solutions,' Langmuir, 10, 3279±3289. 14. Iwasaki Y, Tanaka S, Hara M, Ishihara K, and Nakabayashi N (1997), `Stabilization of liposome attached on polymer surface having phosphorylcholine group,' J Colloid Interface Sci, 192, 432±439. 15. Kadoma Y, Nakabayashi N, Masuhara E, and Yamauchi J (1978), `Synthesis and hemolysis test of the polymer containing phosphorylcholine groups,' Koubunshi Ronbunshu (Jpn J Polym Sci Tech), 35, 423±427. 16. Ishihara K, Ueda T, and Nakabayashi N (1990), `Preparation of phospholipid polymers and their properties as hydrogel membrane,' Polym J, 22, 355±360. 17. Ueda T, Oshida H, Kurita K, Ishihara K, and Nakabayashi N (1992), `Preparation of 2-methacryloyloxyethyl phosphorylcholine copolymers with alkyl methacrylates and their blood compatibility,' Polym J, 24, 1259±1369 (1992). 18. Ishihara K, Tsuji T, Kurosaki K, and Nakabayashi N (1994), `Hemocompatibility on graft copolymers composed of poly(2-methacryloyloxyethyl phosphorylcholine) side chain and poly(n-butyl methacrylate) backbone,' J Biomed Mater Res, 28, 225± 232. 19. Ishihara K, Inoue H, Kurita K, and Nakabayashi N (1994), `Selective adhesion of platelets on a polyion complex composed of phospholipid polymers containing sulfonate groups and quaternary ammonium groups,' J Biomed Mater Res, 28, 1347±1355. 20. Ishihara K, Oshida H, Endo Y, Ueda T, Watanabe A and Nakabayashi N (1992), `Hemocompatibility of human whole blood on polymers with a phospholipid polar group and its mechanism,' J Biomed Mater Res, 26, 1543±1552. 21. Ishihara K, Ziats N P, Tierney B P, Nakabayashi N, and Anderson J M (1991), `Protein adsorption from human plasma is reduced on phospholipid polymer,' J Biomed Mater Res, 25, 1397±1407. 22. Ishihara K, Ueda T, Saito N, Kurita K, and Nakabayashi N (1991), `Suppression of protein adsorption and denaturation on polymer surface with phospholipid polar group,' Seitai Zairyou (J J Soc Biomat), 9, 25±31. 23. Ishihara K, Nomura H, Mihara T, Kurita K, Iwasaki Y, and Nakabayashi N (1997), `Why do phospholipid polymers reduce protein adsorption?' J Biomed Mater Res, 39, 323±330. 24. Lu D R, Lee S J and Park K (1991), `Calculation of solvation interaction energies for

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protein adsorption on polymer surfaces,' J Biomater Sci Polymer Ed, 3, 127±147. 25. Tsuruta T (1996), `Contemporary topics in polymeric materials for biomedical applications,' Adv Polym Sci, 126, 1±51. 26. Kitano H, Sudo K, Ichikawa K, Ide M, Ishihara K (2000), `Raman spectroscopic study on the structure of water in aqueous polyelectrolyte solutions,' J Phys Chem B, 104, 11425±11429. 27. Kitano H, Imai M, Mori T, Gemmei-Ide M, Yokoyama Y, Ishihara K (2003), `Structure of water in the vicinity of phospholipid analogue copolymers as studied by vibrational spectroscopy,' Langmuir 19, 10260±10266. 28. Ishihara K (2000), `Bioinspired phospholipid polymer biomaterials for making high performance artificial organs,' Sci Tech Adv Mater, 1, 131±138. 29. Nakabayashi N and Pashley D H (1998), Hybridization of dental Hard Tissues, Tokyo, Quintessence publishing Co., Ltd. 30. Nakabayashi N, Kojima K, Masuhara E (1982), `The promotion of adhesion by the infiltration of monomers into tooth substrates,' J Biomed Mater Res, 16, 265±73.

15

Blood flow dynamics and surface interactions W V A N O E V E R E N , University of Groningen, The Netherlands

15.1 Clinical application and problems of medical devices in contact with blood Application of biomaterials in direct blood contact results in activation of the blood coagulation system and in an inflammatory reaction. These responses of blood are due to the natural response of the host defence mechanism against foreign surfaces. Inadequate control by natural inhibitors results in pathological processes, such as microthrombi generation or thrombosis, bleeding complications, haemodynamic instability, fever, edema, and organ damage. These adverse events become manifest during prolonged and intensive foreign material contact with vascular implants and extracorporeal blood circulation.1-7 Medical Device Alert in the UK has shown that the number of adverse events induced by CE-marked products is increasing (Fig. 15.1). More than one quarter of those adverse events is due to blood contacting devices. FDA reports indicate a doubled or even tripled increase of thrombotic incidences of implants from 1998 to 2003 (Fig. 15.2). It is hard to speculate on a direct relation with haemocompatibility, but consensus exists about the important role of poor haemocompatibility in direct and sustained adverse reactions. Some of the most studied effects of biomaterials in contact with blood will be discussed in this section.

15.1.1 Small implants in the blood vascular system: stents and vascular grafts After balloon dilatation of narrowed small diameter arteries, stents are frequently applied to maintain the lumen open. The basis material of stents is a metal, such as stainless steel, tantalum, nitinol, based on the mechanical properties of these metals to support the vascular wall with minimal occlusion of side branch capillaries. The occurrence of in-stent subacute thrombosis has been reduced, but still remain as a worrying complication because of its strong impact on short-term mortality.8 Due to the refinement of adjunctive antiplatelet

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15.1 Reports of incidents with medical devices in the UK indicate a significant increase of CE marked products, partly due to an increased use of such CE products, but also indicating the admission to the market of malfunctioning medical devices. Adverse events in circulating blood often cause serious health problems, which may explain the high frequency (28%) of blood contacting devices in this figure (adapted from presentation by Tony Sant, Manager Medical Devices Agency, MDA liaison officer conference 6 March 2002).

treatment9 and the establishment of the most appropriate ways for stent utilisation,10 reduced thrombosis was observed around the millennium change. However, an increase of thrombotic incidents has been observed since, due to increased stent utilisation, and also related to new types of bioactive coatings, aiming for reduction of late reappearance of a coronary stenosis at the site of intervention (restenosis). In-stent subacute thrombosis and restenosis arise partly from the same source. The immediate local deposition of platelets and leukocytes onto the bare material surface can lead to abrupt lumen obliteration and enhances the release of cytokines and growth factors, thus promoting intimal hyperplasia of smooth muscle cells into the vessel wall.11 Intimal

15.2 Thrombotic events are one of the most obvious side effects of implants in the vascular system. According to the FDA (USA) the number of thrombotic events reduced in three types of devices in the years 1996 to 2000. This can be explained by the use of aggressive antithrombotic drugs (on the cost of increased bleeding complications). However, new types of catheters and stents have induced a significant increase of thrombosis in recent years (adapted from: MAUDE database search, Center for devices and radiological health, FDA).

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hyperplasia is an important reason for late stent or graft failure. It is recognised as intrinsic obstructive lesion after many forms of arterial wall injury, for instance, in coronary and superficial femoral artery angioplasty, endarterectomy, arteriovenous fistulae for haemodialyis and homograft transplanted organs.12 Intimal hyperplasia occurs physiologically in closure of the ductus arteriosus after birth.13 Implantation of a medical device into the arterial circulation leads to endothelial denudation, which is immediately followed by deposition of platelets and leucocytes. Besides damage to the endothelium, high arterial pressure and flow cause damage to the medial layer of the blood vessel. Approximately six hours after implantation leucocytes infiltrate the vessel media. This medial layer is to a certain extent damaged resulting in the death of smooth muscle cells. Both dead and injured endothelium and medial smooth muscle cells are able to release growth factors. One of these released growth factors is basic fibroblast growth factor (bFGF) which stimulates the proliferation of endothelial cells and of smooth muscle cells.14 Besides proliferation, the synthetic smooth muscle cells produce extracellular matrix resulting in medial thickening.15 Another important growth factor is platelet derived growth factor.16 Particularly biomaterial implants with poor haemocompatibility and poor rheology will cause platelet activation and subsequent release of platelet derived growth factor. The triggers for the formation of intimal hyperplasia that have been further defined are injury, circulating blood components, and haemodynamics.17±19

15.1.2 Pharmacologic treatment for heart valves, coronary stents and vascular prostheses The most common platelet aggregation inhibitor is aspirin. Aspirin and other non-steroid anti-inflammatory drugs irreversibly acetylate a serine residue in the active site of cyclo oxygenase blocking the formation of thromboxane. Aspirin has no capacity to block the release of platelet derived growth factor nor the capacity to block the first wave of ADP induced platelet aggregation,20 so the effect of platelet aggregation inhibitors is theoretically low. In experimental models of vein grafting, conflicting results are present on the reduction of intimal hyperplasia using platelet aggregation inhibitors.21,22 The currently recommended antiplatelet treatment after stenting (Aspirin + Ticlopidine or Clopidogrel) has been standardised after experimental and clinical observations with the first generation of stents, and since then universally applied. However, it has to be proven that this combination of antiplatelet drugs is suitable for any procedure. Besides, it seems likely that, in particular situations (optimal stent expansion, less device-induced thrombogenicity), patients can profit by a sole, though adequate, antiplatelet treatment.23 Furthermore, the population of cardiopathic patients submitted to percutaneous interventions

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evolves continuously and dramatically fast. Lately stenting procedures have been safely and effectively implemented in the treatment of patients suffering from acute coronary syndromes (unstable angina, acute myocardial infarction). Many pivotal trials have demonstrated that, in such patients, the blockade of activated platelet GPIIbIIIa receptors positively influences the short- and longterm outcome after stenting, by reducing the risk of subacute thrombosis, by optimising the blood flow in the tributary microcirculation, and by possibly preventing the development of restenosis.24 Also patients affected by diabetes mellitus, advanced in years, or in whom particularly complex coronary lesions have been detected, seem to benefit from stent usage, more than by PTCA alone.25±27 In 1977, it was discovered that systemic delivery of heparin suppresses the formation of intimal hyperplasia after injury of carotid arteries in rats.28 Later studies revealed that heparin inhibits proliferation and migration of smooth muscle cells probably by interfering with growth factors and independently from its anticoagulant properties.14,29 Systemic administration of heparin has yielded conflicting results with respect to its effect on intimal hyperplasia in experimental vein grafts.30,31 The late 1980s saw the move away from systemic to local therapies. To minimise systemic effects, delivery of a therapeutic agent locally at the time of the operation would be a logical strategy.32 The current coatings applied to stents are mainly directed to counteract intimal hyperplasia by their inhibiting effect on proliferation. The thrombogenicity of these coatings is not studied in detail, but a high incidence of thrombotic occlusions after implantation of stents coated with drugs that affect the vascular tone indicates a potential serious side effect of these inhibitors of proliferation.

15.1.3 Extracorporeal circulation In 2003 almost 700,000 heart operations worldwide were performed with the assistance of extracorporeal circulation (ECC). With respect to the age and physical condition of patients, a shift toward older adults on one hand and younger children (up to 50% infants, half of them new-borns) on the other was observed. The possibilities of operating on these patients successfully can be explained by continuous improvement in the operative techniques, as well as perioperative supervision and mechanical circulatory support systems. The optimisation of the surfaces of ECC devices is of increased importance, both during the operation (heart-lung machine) and after operation with systems providing sustained heart support. Next to the technical perfection of all these systems, it is also particularly important to consider the surfaces presented to the circulating blood. Since its first use in 1953, extracorporeal circulation (ECC) during open heart surgery has developed into a routine procedure. Nevertheless, an insufficient

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haemocompatibility of materials used for ECC devices still remains a problem. The contact between blood and the various artificial surfaces of the extracorporeal system leads post-operatively to a post-pump syndrome, which can escalate into a systemic inflammatory response syndrome (SIRS),33 acute lung failure (ARDS: adult respiratory distress syndrome),34 sepsis, or even multiorgan failure (MOF).35 The causes of these syndromes are multi-factorial; mechanical and chemotactic activation and membrane-damage of the blood cells, dysfunction of cellular immune regulation, and activation of the haemostatic system. The materials used for extracorporeal application include a wide spectrum of polymers, in particular polyethylene, polypropylene, polyvinylchloride, polyester, polystyrene, polyurethane, and silicone. Although these products possess the required physical properties, they display more or less the same disadvantage; an incompatibility with blood and tissues. Through contact with the blood, this incompatibility can provoke a pathophysiological response from the organism, similar to that of traumatic shock. As is well known, in adult patients undergoing a bypass grafting procedure the total blood volume comes into contact with about 3 m2 of these non-physiological surfaces for one to several hours. This extensive contact causes a massive activation of the humoral and cellular defence systems. Such side effects of biomaterials are counteracted in part by coating surfaces to obtain improved biocompatibility or by pharmacological inhibition of the enzymes responsible for consecutive activation of the cascade reactions

15.2 Surface interactions of blood After contact of blood with a material various proteins will be deposited within split seconds. The main proteins adhered to a surface are albumin, fibrinogen and immunoglobulin, based on their high concentrations in blood. After the initial adhesion a continuous exchange with free proteins takes place, that reaches equilibrium after approximately two hours. This results in binding of higher molecular weight proteins. The relatively medium molecular weight protein albumin will be exchanged in part for larger proteins. Next to nonspecific protein deposition, some components of the contact system react specifically with negatively charged surfaces. As soon as the blood comes in contact with a negatively charged surface, Factor XIIa fragments are formed. These fragments then initiate the entire contact system. -Factor XIIa converts prekallikrein into its active form, kallikrein, which generates the vasodilator bradykinin.36 The deposition and conformation of some plasma proteins on the artificial surface, such as Factor XII, fibrinogen, and vitronectin are a significant criterion for further thrombogenicity.37 After deposition to some surfaces fibrinogen leads to a strong adhesion of platelets through platelet glycoprotein receptors (GpIIbIIIa), followed by platelet aggregation, and release of

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15.3 The main products involved in coagulation, including the intrinsic and extrinsic pathway. Although the intrinsic pathway has physiologically almost no meaning, it plays an important role in activation by biomaterials. The intrinsic pathway is initiated by the contact system (XII and kallikrein) or by Factor XI. Deposition of leucocytes results among others in tissue factor activation, whereas activated platelet membranes contribute to coagulation by complex formation of factors IX, VIII and X, V.

procoagulant contents from platelets. Additionally, contact activation induces activation of the coagulation cascade (Fig. 15.3). Fibrin clots and thrombi are cleared by fibrinolysis, the enzymatic process of fibrin fragmentation and platelet release from binding to fibrinogen. The fibrinolytic enzyme plasmin can be formed through either the release of endothelial tissue plasminogen activator (t-PA) or by kallikrein-activated urokinase. The inflammatory reaction is initiated by complement activation. C3b, which is present in small amounts in blood, after adhesion to a negatively charged surface forms a C3 convertase (C3 cleaving enzyme) when not immediately degraded by complement inhibitors. Since foreign surfaces lack complement inhibiting capacity, the complement convertases will amplify the complement reaction by cleavage of new C3 molecules, resulting in C3b generation and its deposition onto the surface. Simultaneously, the smaller C3a fragment is released in plasma and this fragment is often used as a marker of complement activation. Thus, an exponential activation of the complement system takes place after recruitment of the other complement factors (Fig. 15.4). Similarly to C3 convertase C5 convertase will also be formed, with cleavage capacity for C5 into C5a and C5b conversion. The complement has lytic effects on target cells by its end stage components C5±9 (terminal complement complex) and therefore may become harmful for

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15.4 The alternative pathway of the complement system reacts independently of other immune factors on any foreign surface by deposition of factor C3b and subsequent recruitment of other factors, such as Bb. The C3 convertase cleaves new C3 molecules to form C3b and C3a, which enhances deposition of C3b and causes signalling of leucocytes by release of C3a and C5a. Finally, the terminal complement complex, composed of all factors C5b±C9, is formed. This complex may cause host cell lysis.

the patient in contact with an activating device. Moreover, most of the deleterious effects of complement activation are related to the recruitment and activation of leucocytes, such as granulocytes and monocytes. Granulocytes show an upregulation of the adhesion molecules CD 11b and CD 18 with increased adhesion to the surface, release of elastase and superoxide generation, i.e., further propagation of the inflammatory response.38,39

15.2.1 Leucocytes Leucocytes release a number of inflammatory products including chemotactic factors, growth factors, and complement components. A second mechanism involves the production of lysosomal degradation enzymes. Activated leucocytes elaborate several potent proteases capable of degrading collagen and other structural extracellular matrix and extracellular components, for example, basement membranes. Heparanases can remove heparan sulphate proteoglycans from the cell surface and diminish their inhibition on cell proliferation.18 Lastly, leucocytes may also act at sites of endothelial injury through the production of oxygen free radicals. Granulocytes can produce oxygen free radicals capable of injuring remaining viable endothelium leading to an ongoing stimulation of inflammatory injury.14,18

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Inflammatory processes related to biomaterials have been studied extensively in patients undergoing extracorporeal circulation (heart-lung machine or dialysis). During surgery and the early post-operative stage, the extent of the inflammatory response is associated with clinical symptoms such as fever, bleeding and in severe cases organ failure.1,40 This is more pronounced after use of the heart-lung machine, due to its large surface area and corresponding massive activation of the complement system. More recently its was found that use of the heart-lung machine causes biphasic complement activation. The first phase occurs during operation and directly results from the interaction of blood with the extracorporeal circuit. The second phase occurs post-operatively and is characterised by increasing levels of acute phase proteins such as secretory phospholipase A2 (sPLA2) and C-reactive protein (CRP) that contribute to complement activation.45 Conversely, the inflammatory reaction during CPB may contribute to the post-operative generation of sPLA2 and CRP41 and to post-operative morbidity, although it also functions to promote phagocytosis of injured cells and tissue debris.42 Coatings of the extracorporeal circuits have improved biocompatibility, resulting in reduced complement activation and reduced activation of leucocytes during and after bypass surgery.43±46

15.2.2 Platelets Endothelial denudation exposes the subendothelial matrix and leads to platelet adhesion and aggregation. The subendothelium is completely covered by platelets immediately after denudation. Platelet adhesion requires the interaction platelet receptor Gp1b, plasma von Willebrand factor and fibronectin. Platelet aggregation requires fibronectin, von Willebrand factor or vibronectin, and most often platelet receptor GpIIbIIIa. The adhered platelets release adenosine diphosphate and activate the arachidonic acid synthesis pathway to produce thromboxane A2.47 Thromboxane A2 is a potent chemo-attractant and smooth muscle cell mitogen and leads to further platelet recruitment.48 Once activated, platelets release constituents of their granules. Upon activation and during apoptosis platelets and other cells bud off small parts of their plasma membrane, the so-called microparticles (MP). Extensive in-vitro studies have been reported on platelet-derived microparticles (PMP).49±52 When platelets are stimulated in vitro with agonists such as a combination of -thrombin and collagen or the complement complex C5b±9, they release large numbers of PMP. PMP possess `platelet factor 3 activity', i.e., they facilitate coagulation via exposure of negatively charged phospholipids, thereby providing binding sites for activated coagulation factors V (factor Va), VIIIa, IXa, XIa53,54 and enabling the formation of tenase- and prothrombinase complexes (Fig. 15.3).55 Increased numbers of PMP have been reported in the circulation of patients undergoing cardiopulmonary bypass (CPB), and those suffering from acute coronary ischaemia,56-58 and after plasmapheresis,59 which

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has been associated with a thromboembolic tendency. Also in the pericardial fluid of patients undergoing CPB surgery elevated numbers of PMP have been found.60

15.3 Role of blood cells during flow: rolling of cells, effect of concentration of erythrocytes, expression of adhesive cell receptors The rheological properties of blood influence adhesion of platelets, capture and rolling adhesion of leukocytes as well as their margination in the bloodstream.61 Increasing erythrocyte aggregation correlates with increasing leukocyte adhesion and with more slow-flowing leukocytes near the wall. Thus flowing erythrocytes promote leukocyte adhesion, either by causing margination of leukoctes or by initiating and stabilising the attachment that follows. During cardiopulmonary bypass (CPB) a number of non-physiological events take place including haemodilution, hypothermia, and non-pulsatile flow. As a consequence of these events rheology changes and blood flow may be stimulated to shunt from a less favoured organ to preserve a more vital organ.62 At the onset of CPB the prime solution of the extracorporeal circuitry (ECC) mixes with the patient's blood volume. Dilution of corpuscular blood cells, plasma proteins, a reduction of the colloid osmotic pressure (COP) and oxygen transport capacity takes place. To overcome reduced viscocity by haemodilution, to meet the metabolic demands and protect the tissue from ischemia, moderate hypothermia (28 C) is used. In many institutions the target percentage of red blood cell mass in whole blood (haematocrit) during CPB is 22% to reduce transfusion requirements during open heart surgery.63 If possible, autologous blood is predonated and replaced by an equal volume of an artificial colloid to reach the predicted haematocrit. However, in a study where prime volume was reduced during CPB, the haematocrit in the reduced prime group was significantly higher than in the full prime group, resulting in a reduction in allogeneic blood use.64 Moreover, markers of organ ischaemia are reduced when the haematocrit is kept to a more physiological level. During CPB, specifically in non-pulsatile mode, splanchnic vasoconstriction and hypo-perfusion occur, leading to intestinal mucosal ischaemia, and subsequently to the release of endotoxins from the gut into the circulation.65,66 Circulating endotoxins seem to be one of the most potent factors in CPB that contribute to post-operative morbidity by producing an inflammatory reaction.67 Although the more vital organs such as the kidneys show no adverse effects in the course of CPB,68 during rewarming inadequate perfusion and subsequently potential damage to the superficial cortex may occur.69 Ranucci,70 clearly demonstrated that a haematocrit value below 25% is the primary risk factor in developing severe renal dysfunction. It is suggested that an optimal haematocrit during CPB is not necessarily a lower haematocrit, because it should lead to

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sufficient overall tissue oxygenation and reduce any preferential blood flows to other vital organs, whilst at the same time limiting the requirements for blood transfusion.

15.4 Biomaterial surface characteristics in relation to haemocompatibility and clinical applications A very important requirement for biomaterials used for temporary support of organs or as permanent implants in the human body is minimal generation of thrombosis. Adhesion and activation of platelets to biomaterials surfaces is an important step in thrombosis and is governed, in part, by surface energy and wettability of the biomaterial surface.71 Prior to adhesion of platelets, plasma proteins like fibrinogen and fibronectin adsorb,72 and the composition of the adsorbed plasma proteins relates to the wettability of the biomaterial surface.73,74 Adhesion can be controlled by adjusting the surface properties ± especially surface energy ± of the material involved. Long-term implantation of totally artificial hearts is one of the most compelling proofs of the bioengineering utility of surface energy modification to minimise biological adhesion. These pumps, and the related intra-aortic balloons and left ventricular assist devices, do not accumulate blood clots or thrombotic masses during their contact with blood.

15.4.1 Wettability The effects of wettability on deposition of thrombotic material can be studied by means of wettability gradients. These are made by a gradual change in chemistry along their length, resulting in a varying wettability. Therefore, they are excellent tools for studying the biological effect of this property in one systematic experiment. These wettability gradients have been employed to study various biological phenomena, including protein adsorption,75,76 and cellular adhesion.77±79 Recently, a new, simple method to prepare wettability gradients on polymers by means of glow discharge by partly shielding the material with an aluminum cover has been published.80 Such prepared wettability gradients on polyethylene extended over 4 to 5 cm, and their steepness could be controlled by adjusting the height of the aluminum cover above the polyethylene surface and the duration of the treatment. Consequently, shielded gas plasma produced wettability gradients on polyethylene are very suitable to study biological interactions, because they extend over appreciable lengths and the gradients are relatively stable. It has been hypothesised, that shielded gas plasma-produced wettability gradients are a result of hydrophilic groups on a hydrophobic base.81 Albumin, fibrinogen and immunoglobulin G are the most prevalent proteins in blood plasma82 and their adsorption along the length of a shielded gas plasma-

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produced wettability gradient on polyethylene increased with the distance from the hydrophobic end. Since most implanted devices are in contact with blood under flow conditions, it is also relevant to study the effect of wettability of a device surface on the adhesion and activation of platelets under conditions of flow, especially because high flow may trigger the adhesion and aggregation of platelets.83,84 In vitro, shear force induced platelet activation and adhesion to collagen occurs within 2 s, with half of the number of reacting platelets adhering within 240 ms.85 This is important for efficient haemostasis under flow conditions and in the contact of blood with medical devices. Generally, more platelets adhere to the hydrophilic than to the hydrophobic end of a gradient, while flow promotes platelet adhesion evidently through increased convective mass transport.86 Moreover, attachment is known to be stimulated by shear stress, which causes haemostasis under arterial flow conditions.87±90 A moderate flow and shear stress (0.8 N/m2) generated the most pronounced difference in platelet adhesion along the gradient surface. However, when the flow was further increased to simulate the conditions of coronary arteries at 3.2 N/m2, platelet numbers at the hydrophilic end were significantly reduced as compared with the hydrophobic end. These results strongly suggest detachment of platelets from hydrophilic surfaces. Within 15 min at high shear force platelet deposition on hydrophilic surfaces will be limited, probably by detachment after initial adhesion. Such effects can be explained by the small contact area of platelets with hydrophilic surfaces.78 Furthermore, the platelets attached to hydrophilic surfaces remain spherical and extend deeper in the blood flow, thereby experiencing higher shear forces. In contrast, platelets on hydrophobic materials can withstand high shear forces due to strong contact and complete spreading of platelets. When examined with scanning electron microscopy, the platelets on the hydrophobic end of a gradient surface were indeed more extended like a pancake than on a hydrophilic end.91 Obviously, the gradient surfaces at the hydrophobic end reacted similarly with platelets under a high shear stress of 3.2 N/m2 as under a low shear stress of 0.8 N/m2, which indicates a relatively strong binding of platelets on the hydrophobic surface. It can be concluded that under conditions of arterial flow, especially at the hydrophilic end of the gradient surface, fewer platelets adhere after 15 min due to detachment.92 It is hypothesised that hydrophilic device surfaces exposed to flowing blood in the human body and under high shear conditions, are less likely to accumulate platelets than hydrophobic surfaces.

15.4.2 Roughness The influence of biomaterial roughness on thrombogenicity is not clear, since various studies show apparently conflicting results. Bailly compared angiographic catheters with different surface roughness by their tendency to become occluded. They concluded that the most thrombogenic material was the

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smoothest, whereas surface chemistry (polyethylene versus polyamide) contributed to a lesser extent to thrombogenicity.93 Zingg et al. found that increased roughness caused a decrease in platelet adhesion on hydrophilic surfaces and an increase on hydrophobic surfaces. These results were obtained when flow conditions were applied. During static test conditions no differences between smooth and rough surfaces were found.94,95 One explanation for different observations is the higher extent of thrombogenicity at smooth surfaces, whereas the degree of thrombus adhesion is higher at rough surfaces.96 In a more detailed study it was observed that roughness due to titanium crystals appeared to initiate more activation of the clotting cascade, but less platelet adhesion.97 These different effects of two important factors of thrombus formation in conjunction with the variability induced by various flow and shear stress conditions and wettability may explain the conflicting results regarding the thrombogenicity of biomaterials.

15.5 Haemocompatibility of metals, ceramics and polymers Due to their mechanical and radio-opaque properties metals are frequently used for manufacturing implant devices and as part of devices used for invasive procedures for diagnostic and therapeutic purposes. Often, these implants are in direct contact with blood, e.g., as stents, heart valves and catheter tips. The blood compatibility data of metals are relatively scarce, which is possibly due to the historically accepted application of metals as medical devices, or to the indispensable physical characteristics of metals. In reviewing the literature, some frequently used metals even appear to have prothrombotic properties which, if alternatives were available, would lead to refusal of its use in bloodcontacting devices or implants. No thorough comparisons between metals can be made, since the reported studies have evaluated the blood compatibility of different materials and no consistent reference materials were included. Furthermore, most studies report only limited blood compatibility tests. Apparently, the possible induction of an inflammatory reaction, initiated by complement activation or granulocyte activation is not frequently tested, although it is an important contributor to intimal hyperplasia.98 It can be concluded that the more noble metals appear less blood compatible than the oxidised titanium and aluminium metals (ceramics) and silicon carbon products. Most bare metals had a poor blood compatibility in direct comparison to polymers, which was most often tested after polymer coating of the metals, indicating that the mechanical properties of metals are still considered essential for stent or valve construction. A thorough evaluation of the blood compatibility of metals is warranted to quantify their thrombotic and inflammatory properties. Some of the most frequently used metals are now discussed.

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15.5.1 Stainless steel Stainless steel (316L) is the most commonly used metal for endovascular devices. Its mechanical properties significantly contribute to its applicability, but the blood compatibility results also appear better than those of some other metals. For instance, stainless steel stents are more blood compatible than tantalum stents.99 However, stainless steel can also be further optimised, since several studies showed that polymer coating of stainless steel stents reduced deposition of platelets and thrombus mass by more than 60%.100,101 The reported reduction by stainless steel of the partial thromboplastin time by 50% or more, indicating activation of the clotting system, would be considered unacceptable in view of the requirements for newly developed biomaterials.

15.5.2 Tantalum After stainless steel, tantalum was introduced as the metal for the construction of stents. Due to tantalum's high radiopacity, implantation of tantalum stents is greatly facilitated. Initial studies showed similar blood compatibility for tantalum and stainless steel,102 although later studies indicated that stainless steel possesses better blood compatibility.99 Clinical studies indicated that a high incidence of thrombotic complications could occur after tantalum stent implantation if anticoagulation and anti-platelet therapy was insufficient.103 Also, post-stent antithrombotic therapy was required, including both anticoagulants and platelet inhibitors or Ticlopidine plus Aspirin.104 Polymer coating of tantalum stents with polyurethane or parylene reduced the deposition of platelets by 5 to 50% relative to platelet deposition on uncoated stents.105

15.5.3 Titanium In the human body, titanium exists only for a short period of time in its unmodified form, and relevant blood compatibility data are therefore obtained with titanium nitride or titanium oxide. Titanium oxide appears to reduce fibrinogen deposition due to its semi-conductive nature. This effect is explained by the similar electronic structures of fibrinogen and titanium.106 In a comparative study with low-temperature isotropic pyrolytic carbon (LTI carbon), not only reduced deposition of fibrin, but also a 50% reduction in microscopically counted platelets was observed with titanium oxide.107 Transvenous inferiorvena-cava filters made of stainless steel, titanium or titanium-nickel all showed approximately 25% early thrombosis in clinical use, measured via ultrasound scanning.108 This incidence of early thrombosis was unexpectedly high, and difficult to reduce with the current devices, since antithrombotic medication is often contra-indicated in patients requiring a vena cava filter. Titanium nitride has been tested for its blood compatibility with regard to

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leukocyte adhesion, and appears to retain no leukocytes.109 Additionally, platelet retention was as low as with silicone elastomer, but comparisons with more blood compatible materials have not been made.110 In-vivo experiments with titanium nitride heart-valves in sheep showed some deposition of fibrin and platelets.111

15.5.4 Nitinol Nickel-titanium alloy (Nitinol) has attracted special attention due to its shape memory function. It must be noted that Nitinol has an outer surface of titanium (oxide), whereas nickel is not exposed to blood. Therefore, blood compatibility characteristics are expected to be rather similar to those of titanium oxide. Based on the hypothesis that a semi-conductor prohibits fibrin and platelet deposition, Nitinol is expected to be thromboresistant, unless its semi-conductive nature is lost in the alloy. In a clinical study with Nitinol intravascular clot filters, the effects on the clotting system and on platelet adhesion were shown to be similar to those induced by stainless steel.112 An experimental study with stented rabbits showed significantly more thrombus formation on stainless steel than on Nitinol.113 However, grafting of polyethylene oxide (PEO) on Nitinol reduced the fibrinogen adsorption by as much as 99%, and significantly reduced platelet adhesion, which once more shows the superior thromboresistant effects of polymers as compared to those of metals like Nitinol.114 Further evidence that Nitinol, too, can only be safely implanted during antithrombotic treatment was provided in experiments that included the use of platelet inhibitors Aspirin and Copidrogel in a porcine stent model. Combined treatment with these inhibitors reduced stent thrombosis by 95±98%.115 An effective coating such as PEO could thus limit the use of systemic treatment by medication.

15.6 Biological surface treatment to improve haemocompatibility Immediate stimulation of platelets during contact with non-albumin coated extracorporeal circuits has been observed by the release of platelet granules. After this first contact no further substantial platelet stimulation occurred. Albumin has been described to inhibit this release reaction of platelets and to inhibit platelet aggregation.116,117 The first pass effect can be reduced by initial adsorption of albumin to the biomaterial surface. Fibrin deposition and platelet GpIIIa receptor binding was also reduced on tubing of an extracorporeal system after pre-coating with albumin. It appeared that a low concentration of albumin was as effective as a high concentration on reducing platelet activation and on reducing deposition onto the tubing. Since albumin priming seems to exert an inhibiting effect only in the first period of extracorporeal circulation, the first

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pass effect on untreated biomaterials seems to exert the most pronounced haemocompatibility problem. However, the initial albumin coating will be replaced by other proteins during prolonged blood contact as a result of the Vroman effect.118,119 A further improvement of the extracorporeal circuit is supposed to be related to more irreversible surface treatment. Dependent on the treatment, the surface may be modified to induce less thrombogenic or inflammatory reactions. Heparin, cell membrane phospolipids, and block copolymer coatings are often used on a number of blood contacting devices.120

15.6.1 Heparin coating The attempt to coat artificial surfaces with heparin, a natural anticoagulant, illustrates the first step to improve haemocompatibility. Gott reported in 1963 on heparin coating of synthetic materials, which had been pre-treated with colloidal graphite.121 Larm122 presented a heparin-coating method in 1983, which is still the most stable and most effective for long-term use. Their method involving covalent binding by the technique of end-point immobilisation, did not adversely affect the heparin active structure and thus produced a bioactive surface. After the first heparin-coated extracorporeal circuits became available in the last half of the 1980s, its haemocompatibility was shown in in vitro systems, animal models, and patient studies.123±125 It can be concluded that heparin coated circuits can cause a reduction of activation of the contact phase, complement system activation, inflammation, and pulmonary complications.126±128 The reduced thrombogenicity of the heparin-coated surfaces was thought to be attributable to the inhibition of thrombin by catalysing the binding to antithrombin III. However, more recent data show that the advantage of heparin coating lies much more in the reduced, or selective, adhesion of plasma proteins. This leads to a faster formation of a blood-friendly secondary layer and prevents a further denaturation and hence activation of the adhered proteins and blood cells. To minimise adsorption of proteins and attachment of cells, next to the heparin effect, Trillium Bio-passive Surface has been developed. This technique works with water-soluble synthetic polymers that are immobilised in two superficial layers. The first polymer is a primer and is based on polyethyleneimine that is modified hydrophobically to allow for strong binding to artificial materials of the medical device. The second layer, containing sulfonate groups, polyethylene-oxide chains, and heparin, is covalently bound to the primer and results in an insoluble surface coating. This surface prevents the adhesion of blood cells and plasma proteins during contact with blood by its negative charge, heparin and PEO chains.129±132 The first experiments with this technique indicated that the treated synthetic materials prevented the activation of the complement system, the contact system

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as well as platelet and leukocyte activation. The TBS coating intervened in the initial phases of the blood-material interactions and thereby prevented further activation at a very early stage in the cascade reaction. ECC systems with this coating did not, however, significantly improve the clinical outcome of systemic heparinised adult patients undergoing cardiac surgery.133

15.6.2 Phosphorylcholine coating Phospholipids that mimic natural cell membranes are currently used as a coating substance. Phosphorylcholine-containing lipids dominate in the outer membrane of the cell membrane bilayer and these appear to possess strong antithrombotic properties.134 One has succeeded in coupling synthetic methacryloylphosphorylcholine/lauryl-methacrylate copolymers to metal and synthetic surfaces. The term `biomembrane mimicry' arose for phosphorylcholine-coated foreign surfaces.135 In-vitro experiments and animal tests have shown that phosphorylcholine-coated artificial polymers possess outstanding thrombogenic resistance and display only minimal adhesion of plasma proteins and platelets.136,137 This coating technique has been offered among others for contact lenses, stents and extracorporeal circulation devices. Since coagulation in infants is more delicate than in adults by the reduced availability of inhibitors, this antithrombogenic coating was anticipated to be most profitable for paediatric cardiopulmonary bypass. Although the literature shows an improved biocompatibility in adult surgery when using coatings,138 the thrombogenic and inflammatory response is usally mild in routine adult surgery which makes it difficult to demonstrate differences in post-operative clinical response. Small infants are much more vulnerable to the adverse effects of cardiopulmonary bypass due to the relatively high priming volume and large blood-foreign material surface in combination with the immaturity of several organ systems. A clinical study on small infants showed a reduced thrombogenic and inflammatory response after the use of phosphorylcholine coating, by reduced progression of beta thromboglobulin release and thromboxane production, both related to platelet activation, and by reduced complement activation. While the surface characteristics improved, the coating did not affect the gas transfer properties of the hollow fibre membranes. The characteristic feature of biological membranes is their functional and compositional lipid asymmetry, which has been described in several cell types and is thought to stem from the requirement for biological membranes to have asymmetric protein distributions across the bilayer. In all of the cells for which lipid compositional asymmetry has been described, negatively charged phospholipids are found predominantly on the inner cytoplasmatic side of the membrane, while the neutral zwitterionic phosphorylcholine-containing antithrombotic lipids predominate in the outer membrane leaflet. Negatively charged phospholipids are thrombogenic and it has been proposed that this

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membrane asymmetry may serve the biological purpose in the maintenance of the delicate balance between haemostasis and thrombosis. In-vitro experiments, in which various phospholipid coatings were applied to surfaces, showed a very high procoagulant activity of negatively charged phospholipids. This is in contrast to phosphorylcholine-containing surfaces that were not active in coagulation tests.139,140 However, inhibition of activation of the clotting system was not observed, which may indicate a merely passive effect of the phosphorylcholine coating towards the clotting system. Previously, heparin coating has been evaluated in paediatric CPB.141,142 As in adult CPB heparin coating reduced complement activation. Surprisingly, also the phosphorylcholine coating appeared to generate less complement activation than the uncoated systems. In-vitro experiments showed decreasing complement activation with increasing surface phoshorylcholine mole fractions,143 suggesting that phosporylcholine is responsible for the reduction. The working mechanism is probably related to lesser activation of the complement protein C5144 and the inhibition of monocyte and macrophage adhesion.145

15.6.3 SMA coating Surface modifying additives (SMA), are mixed with the initial synthetic materials in the production phase, and this technique is therefore not a coating in the usual sense. The copolymer distributes itself in the synthetic materials during the polymerisation process and due to its charge characteristics, moves to the surface of the basis material as it cools. Thus, a new surface of primarily SMA forms. The microscopic structure of the surface of alternating hydrophilic and hydrophobic regions carries a zero net charge, thereby reducing platelet and leukocyte deposition. Tsai et al.146 could prove that SMA surfaces decreased coagulation activation and significantly reduced contact phase and complement activation. Gu et al.147 found better platelet protection in clinical CPB by using SMA treated devices. However, larger clinical studies on routine cardiopulmonary bypass patients showed only minor clinical benefit of SMA treated devices.

15.6.4 PMEA coating Poly-2-methoxyethylacrylate (PMEA) is a hydrophilic polymer coating that minimises the adsorption and denaturation of proteins and blood cells. In various animal and clinical studies, this coating has been proven to reduce blood activation during extracorporeal circulation. When compared with uncoated oxygenators, PMEA-coated oxygenators exhibited increased thrombus resistance with lower inlet pressure and lower thrombocyte consumption. Plasma bradykinin levels and the percentages of activated monocytes in PMEA-coated circuits were significantly lower than those in uncoated circuits during CPB. The

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amount of protein adsorbed on PMEA-coated circuits was significantly lower than that on uncoated circuits (0.30 g/cm2 versus 3.42 /cm2). Almost no IgG, IgM, or C3c/d was detected in proteins adsorbed to the PMEA-coated circuits although these proteins were clearly detected on the surfaces of uncoated circuits.148 A clinical study showed no significant differences between heparin-coated and PMEA-coated groups in the plasma concentrations of inflammatory markers, or markers of clotting. Clinical variables did not differ significantly between the groups. It was concluded that PMEA-coated CPB circuits are as biocompatible as heparin-coated CPB circuits and prevent post-operative organ dysfunction in patients undergoing elective coronary artery bypass grafting with CPB.149 The cost-effectiveness ratio seems favourable for PMEA-coated circuits.150

15.7 ISO 10993 requirements for testing of medical devices: simulation of clinical application including flow, blood composition, anticoagulants In December 2002 the revised ISO 10993-Part 4 standard (Biological evaluation of medical devices ± Selection of tests for interactions with blood) was published.151 Ten years of discussions, conferences, writing and implementation of new insights preceded the revision of this standard that deals with the reaction of blood to medical devices. The standard is applicable to external communicating devices, either with an indirect blood path (e.g. blood collection devices, storage systems) or in direct contact with circulating blood (e.g. catheters, extracorporeal circulation systems), and implant devices (stents, heart valves, grafts). Testing should be performed for five categories, based on primary processes: thrombosis, coagulation, platelets, haematology and complement. In this system all relevant aspects of blood activation are taken into consideration, but, and this is most important, testing should simulate clinical conditions as much as possible. The increased use of medical devices for temporary use or implant in the blood circulation has resulted in increased demand for evaluation of complications brought about by these devices. One important and extremely relevant aspect of testing of medical devices is the condition of blood exposure to the device. Often, blood with clinically inapplicable anticoagulants and under static conditions was incubated with the test device.152±154 Currently anticoagulation and flow conditions must be as similar as possible to the clinical application to achieve relevant test results. Thus, most devices must be tested with heparinised blood under circulating conditions. For some devices, such as stents and catheters this implies high flow through or around the device to obtain relevant shear stress conditions. The major differences observed between cell interaction

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under static and flow conditions has made clear that whole-blood flow models are required for testing haemocompatibility inasmuch as the test device will be used clinically in the blood circulation. Flow models for testing may consist of animal models or in-vitro test systems. Animal models have the disadvantage of being expensive, time consuming and insensitive due to overwhelming shortterm effects of tissue damage. Animals, particularly, are being used to test haemocompatibility of implants. Clearly, human volunteers cannot be used for this purpose. Animal models include the effect of tissue damage by operation and the important antithrombotic effects of endothelial cells. On the other hand the extent of haemocompatibility is obscured by these tissue effects. Moreover, it has been shown that the composition of blood differs considerably between various species, which leads to over- or under-estimation of human blood reactions to biomaterials.155,156 The use of human blood is therefore more relevant to the interpretation of results and offers a more detailed array of test methods, since most available methods are based on human blood components. The use of human blood requires a proper in-vitro circulation model, which is discussed in the next section.

15.8 Test models: static, low flow, arterial flow, pulsatile/laminar flow A key determinant of blood activation and adhesion of cells is wall shear stress; the force exerted by the flow per surface area. In a cylindrical tube this property is easily calculated from  ˆ 32Q=…D3 † where  is the shear stress, Q is the flow rate,  is the viscosity, and D is the tube diameter.157 In configurations that differ from the cylindrical tube, such as just after a bifurcation, wall shear stress has much larger values locally, than at the opposite site.158,159 Growth of intimal thickness is often observed at locations with low shear stress.157,160 Wall shear stress in the normal circulation is rather constant when the equation is applied to blood vessels of various sizes.157 Values are found in the range 10±20 dynes/cm2 (1±2 Pa). Furthermore, blood vessels adapt their diameter as much as possible towards a constant value for shear stress.161 Also, when blood is in contact with biomaterial surfaces, fluid mechanics, and especially the shear stress, have a strong influence on the damage of red cells and platelets. Red cell damage may occur at high shear stress.162,163 Platelets are more easily damaged by shear stress.162,164 Platelet damage is not only influenced by the maximum shear, but also by the duration of the shear force. Only for very short exposure times are platelets able to withstand higher shear stress than red cells.164 From a fluid mechanical point of view, differences in flow situations may therefore lead to different problems with blood. Artificial heart valves may cause problems for red cells due to short duration high shear,165 whereas stents

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in the coronary arteries induce intimal growth at locations of relatively low shear,166 which may be caused by platelet activation in high shear. Neointima formation in stents has been shown to be related to wall shear stress as well.166 In tubing used during dialysis, the high shear rate at the needle may lead to problems for red cells,167 but it should not be disregarded that the wall shear stress of the tubing is the most critical issue for platelet activation. Heart valves, extracorporeal systems and vascular grafts or stents induce relatively high shear forces that may result in platelet activation. Shear stress is a natural activator of platelets. The shear-induced pathway appears to be one of the major pathways of platelet induced haemostasis and thrombosis.90 The sequence of the shear-induced pathway is the binding of von Willebrand Factor (vWF) to the platelet glycoprotein Ib (GpIb) receptor, the expression of activated GpIIbIIIa receptors and release of platelet vWF. Finally, vWF binds to GpIIbIIIa, leading to irreversible adhesion. High shear stress of 120 dynes/cm2 induces immediate expression of GpIIbIIIa receptors and release platelet vWF multimers.168 However, in the presence of platelet activators, such as epinephrine and ADP, shear stresses of 60 dynes/cm2 may synergistically result in platelet aggregation.169 Fibrinogen appears to mediate platelet aggregation efficiently at low shear rates, but not at high shear rates.170 Moreover, resting platelets do not adhere efficiently to fibrinogen-coated surfaces; activation by ADP is required.171 For understanding the effects of platelets during use of biomaterial implants or extracorporeal systems these observations are important. A biomaterial surface is initially mainly coated with albumin, immunoglobulins and fibrinogen from plasma. Adhesion of platelets to fibrinogen onto these surfaces is not as irreversible as to collagen coated with vWF multimers of a damaged blood vessel. Additionally, platelet activators significantly support, or are even required for platelet adhesion to fibrinogen, which means that concomitant tissue or blood damage, such as surgical trauma, occlusive pumps, or poorly designed heart valves, synergistically contribute to platelet activation under shear stress. The wettability of biomaterials relates to the extent of hydrophobicity and hydrophilicity. An increased wettability on a polymer gradient was related to an increased amount of oxygen incorporated in the material surface85 and appeared to correlate with increased protein adhesion, activation of the coagulation system, and increased platelet adhesion.96 A classical in-vitro test model is the Chandler loop,172 which consists of a closed tubing partly filled with air, which circulates the device constantly with an airliquid interface. This method may induce artefacts due to the major forces applied to blood elements and protein denaturation at the air-liquid interface.173±176 Thus, instead of the Chandler a small roller pump closed-loop system was used in the past. This model appeared effective for short-term circulation.177±179 An initial experimental blood circulation model with a roller pump, already refined in previous studies,180 appeared efficient, reliable and cost-effective in

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assessing the haemocompatibility of stents, possibly before their clinical utilisation. However, blood damage induced by the pump limited the exposure of the test object to circulating blood. Since metal stents are commonly of a thrombotic nature, experiments for 15 minutes yielded sufficient information to compare stents, but a less traumatic circulation system is required for testing of low thrombotic materials. The Chandler model induces less blood damage than the roller pump, but has the major disadvantage of continuous blood-air contact and the limitation of blood flow due to the requirement to keep air at the top of the circuit. Since improvement of the model by minimising blood damage may increase sensitivity and permits prolonged blood exposure,181 we constructed a simple mechanical device without air and without a pump to reduce blood damage and activation by the device. Moreover, the new device provided pulsatile flow at a frequency similar to the arterial circulation. This haemobile was compared with the Chandler and roller pump model for intrinsic blood damage and finally for testing of thrombosis induced by arterial stents. Fast screening of the thrombogenicity of stents, catheters, vascular grafts and other small medical devices is now possible. The adjustable flow and shear in the haemobile renders it a model that allows standardised testing of these devices at the cost of low intrinsic blood damage. A limitation of in-vitro models is mainly represented by the absence of an endothelial layer in the circulating system. Throughout the release of (anti)thrombotic components and the expression of adhesion molecules, endothelium has a major role in mediating the interplay between the injured vessel wall and blood cells after coronary stenting,182,183 and lack of this character can somehow alter the likelihood of our experimental representation. Nevertheless, all the other elements depicting the blood-stent phase boundary scene are present in the in-vitro model.

15.9 Conclusion In spite of all the technical improvements made to improve the haemocompatibility of ECC components, a noticeable activation of plasma proteins and corpuscular blood components still exists. The long-range aim remains the creation of an optimally haemocompatible surface (endothelium-like), which blood would no longer recognise as unphysiological and hence would not induce humoral and cellular defences as well as rejection mechanisms against it. With the results of numerous studies on the haemocompatibility of extracorporeal circulation under consideration, one can definitely work on the assumption that heparin-coated devices retard the activation of cellular and humoral mechanisms. Above all, the reduction of a general inflammatory response initiated by extracorporeal circulation represents an important application for heparin-coated devices. Above and beyond the long-term applications, routine heart operations have also markedly begun to utilise heparin-coated

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devices. This trend in improving haemocompatibility with active coatings will assuredly continue in the coming years. Heparin coatings are merely the beginning of improved haemocompatibility for all materials that come into contact with human blood or tissues. Intelligent materials with almost completely physiological surfaces will be at the surgeon's disposal within the next few years. Such materials will be able to mimic endothelial functions and respond to individual changes in the patient's status with the controlled release of adequate pharmacological substances. However, control of the quality of patient treatment and a more individual approach is essential to improve the quality of treatment with medical devices. On-line monitoring of organ damage (point of care testing) is essential in order to apply optimised conditions. The currently used markers of organ dysfunction are too non-specific or too much dependent on an advanced stage of deterioration of the organ. Therefore, newer, sensitive organ damage markers for early organ damage will probably be introduced in routine practice.184±186 Bare metal parts exposed to blood should be avoided, either by coating, oxidation or replacement by haemocompatible polymers. Further, the blood composition in terms of haematocrit and plasma expanders should be optimalised, to ensure proper oxygen delivery to all organs. A new important item is compliance with magnetic resonance imaging (MRI) techniques, which will be a common method of screening for malignancies, inflammatory sites and infarctions. Stainless steel and nickel could particularly raise problems in MRI scans. Finally, thorough haemocompatibility testing should have a prominent place in the certification of blood contacting medical devices, since a poor haemocompatibility has long-lasting negative effects on the whole body through blood transport of activation products and on functional recovery at the site of the implant.

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Cell guidance through surface cues A K V O G T - E I S E L E , Max-Planck Institute for Polymer Research, È U S S E R , Institute for Thin Films Germany, A O F F E N H A and Interfaces, Research Centre JuÈlich, Germany and W K N O L L , Max-Planck Institute for Polymer Research, Germany

16.1 Introduction The guidance of cellular adhesion, migration and network formation in vitro by surface associated cues is of great interest for a number of issues both in basic research and in biotechnological applications: Firstly, processes influenced by cues offered on the substrate, e.g., cell adhesion, migration or morphological alterations can be investigated. This is usually done on surfaces exposing patterns of biologically relevant molecules, such as growth factors or proteins of the extracellular matrix. If cells are seeded, e.g., onto surfaces containing alternating stripes of two different matrix proteins, cellular behavior in choice situations can be studied, monitoring attraction, repulsion or morphological changes along the pattern axis. Secondly, the ability to guide cells and the outgrowth of their processes by surface associated cues is useful for the design and the direction of multicellular assemblies with experimentally defined geometries. Examples of such assemblies are orientated vascular tubes formed by endothelial cells or neuronal networks that interconnect along the geometry of an underlying micropattern. This definition of network architecture can be achieved by presenting a surface composed of adhesive and antiadhesive areas, such that cells are forced to adhere to and connect along the permissive regions, while avoiding the antiadhesive background. Patterned neuronal networks allow the study of signal transduction and signal processing in networks of reduced complexity (since only a few pathways are allowed for cellular connectivity) and of defined geometry (as the connectivity pattern is experimentally determined). Thirdly, the experimental manipulation of cellular outgrowth and connectivity on an artificial surface is a promising tool in biotechnology. One potential application lies in the design of implants and prostheses containing grafted cells, whose interaction with the surrounding tissue of the patient may be controlled by providing defined pathways and interfaces for circuit formation. In addition, cell patterning in combination with microfluidic devices offers attractive possibilities for the creation of novel biosensors. It allows locating

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single cells to defined sites on the surface, positioning them on areas exposed to a stream of applied test substances while others are tethered to unexposed areas. In such arrangements, the impact that drug application to single cells has on the behavior of an entire network can be investigated. A number of different techniques have been applied for surface patterning, which we will briefly review with regard to the guidance of neuronal adhesion and network formation. Methods for surface patterning can be roughly divided into two types. On the one hand, cells have been shown to be guided by topographical cues consisting of shapes and textures in the substrate. On the other hand, differences in the chemical properties of confined surface areas can be employed to direct cell-substrate interactions. Figure 16.1 shows a single neuron following defined pathways of a `cell friendly' pattern while avoiding the repellant background material. However, it is important to bear in mind that the effects of surface chemistry and surface topography are not two entirely separate issues. Differential deposition of culture medium components and/or cell derived matrix proteins on surface features may add (bio)chemical differences to topographical structures, while the chemical modification of a surface also has consequences for its physical qualities such as topography, roughness or elastic properties. In the following subsections, we will briefly compare topographical and chemical patterning methods and give an overview over their history in the application of in vitro neuronal cell patterning. In the second section of this short review, we will summarize general principles for surface modifications by chemical grafting and the involved chemistries, and in section three we discuss

16.1 SEM image of single neurons isolated from the cortex of embryonic rats (18 days gestation) on a micropatterned surface. Patterning was performed by microcontact printing using a blend of extracellular matrix proteins and a background of polystyrene.

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the techniques most commonly used in order to achieve patterned cell growth. Finally, in section four we will focus on the biology of neuronal networks and the impact of different types of surface modifications on cellular physiology as well as recent advances in the geometrical design of neuronal networks.

16.1.1 Topographical patterning The employment of topographical patterning techniques started in the early 1960s (Rosenberg, 1963) and 1970s. Cells were seeded onto planar substrates with etched or scribed grooves to study their alignment with the surface structure. With increasing evidence that curvature was the most important parameter for cell guidance, more groups began to examine the effects of varying groove depth, width, and spacing. In the 1980s, lithography was used to microfabricate grooved surfaces, in particular, by utilizing anisotropic etching of silicon wafers. For further details about concepts, materials, surface structures, and possible cellular and biomolecular mechanisms for topographically patterning we refer to reviews by Curtis and Wilkinson (1997, 1998) and Jung and coworkers (Jung et al., 2001). In the late 1980s and early 1990s, the group around Adam Curtis and Chris Wilkinson in Glasgow started to systematically study the relative effects of groove depth and pitch on cell guidance, applying ultrafine structured quartz and silicon surfaces produced by electron beam lithography (Clark et al., 1990). Later, surfaces containing both chemical and topographical cues were applied in order to determine their relative impact on cell guidance if conflicting directions were imposed by the different systems (Britland et al., 1996). The authors determined the critical depth of topographical features required to override guidance through adhesive stripes of aminosilanes; in addition, synergistic effects were reported if chemical and topographical cues were aligned. A further approach to patterning using topographic cues applies pits and grooves sufficiently deep to trap individual cell bodies and outgrowing neurites (Fig. 16.2). This tactic has been used successfully by the group of Peter Fromherz, who achieved the immobilization of snail neurons on a polyester microstructure which created a `fence' around single cells thus confining them to defined surface spots (Merz, 2002).

16.1.2 Chemical patterning First descriptions of chemical cell patterning were presented in the mid 1960s when Carter and his group discovered that fibroblast adhered preferentially to palladium islands evaporated on a polyacetate surface (Carter, 1965). In 1975, Letourneau applied this method to study the alignment of chick dorsal root ganglion neurons with palladium spots on polymeric substrates (Letourneau, 1975). He could demonstrate that the cells preferentially adhered to the metal

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16.2 Scanning electron micrograph of a topographical structure for the defined formation of neuronal networks. The structure was prepared with SU-8 photoresist. The pit has a diameter of 20 m while the width of the groves is c. 10 m.

regions if these were evaporated onto tissue culture plastic substrates but behaved quite indifferently if the palladium was surrounded by polyornithine. This work showed that differences (`contrast') between adjacent regions were necessary to obtain cell patterning. In 1988, Kleinfeld, Kahler and Hockberger used photolithographic techniques to pattern silicon surfaces with small organic molecules in order to direct the adhesion and outgrowth of neurons (Kleinfeld et al., 1988). This study presented a milestone in the field of neuronal cell patterning, as it systematically analyzed the impact of defined chemical groups on cell attachment. It was found that diamines and triamines, but not monoamines, are supportive of cell attachment while surface bound alkanechains have a strong repulsive effect. Similar to photolithography, organic thin films can be patterned by photochemical reactions. The group of Bruce Wheeler applied selective laser ablation to create high resolution grids of polylysine with varying line width, intersection distance and nodal diameter onto which rat hippocampal neurons were seeded (Corey et al., 1991). Photoablation was also used by other groups to pattern ultrathin polymer layers in order to control the adsorption of proteins and the adhesion and spatial orientation of neuronal cells on surfaces (Bohanon et al., 1996). In the middle of the 1990s the Aebischer group studied the impact of surface bound electrical charges on neuronal attachment and differentiation using patterned polymeric substrates (Valentini et al., 1993; Ranieri et al., 1994). An additional approach to surface patterning by chemical methods is presented by the deposition of organic compounds through plasma polymerization. The method tolerates a variety of different chemistries and permits the creation

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of surface features in the m range if parts of the surface are shadowed by a TEM (transmission electron microscope) mask during the deposition process. The suitability of this method for patterned cell cultures has been demonstrated recently (Bullett, 2001; Mitchell, 2002). Towards the end of the 1980s, evidence accumulated that the extracellular environment profoundly influences cellular differentiation in vitro as well as in vivo (Kleinman et al., 1987). Researchers, therefore, increasingly turned to physiologically occurring matrix proteins rather than synthetic substrates for their cell cultures, the simplest method for surface coating being protein adsorption from solution. In 1986, the Arginine-Glycine-Aspartate (RGD) motif was identified as an epitope shared by many matrix proteins that is central to cell-matrix interactions through binding to integrins, a family of adhesion molecules (Ruoslahti and Pierschbacher, 1986, 1987). Shortly afterwards, a number of additional motifs responsible for cell-matrix interactions were identified (Graf et al., 1987; Kleinman et al., 1990). Following the trend in using physiologically occurring adhesive molecules for cell cultures, both entire proteins and isolated recognition peptides were also employed for micropatterning. In 1987, Hammarback selectively destroyed surface adsorbed laminin in defined positions by UV irradiation through a grid mask. The retention of native and therefore adhesion promoting protein in the areas protected by the mask was demonstrated by the guidance of cells along the narrow stripes of protected laminin (Hammarback, 1985). Several years later, the group of Peter Fromherz successfully applied photolithography to pattern ECM proteins for the guidance of leech neurons in culture (Fromherz et al., 1991; Fromherz and Schaden, 1994). Similarly, adhesion promoting peptides were patterned by laser-lithography (Matsuzawa et al., 2000) or microcontact printing (Scholl et al., 2000). In both studies, cell attachment to the permissive areas and neurite outgrowth along the desired pathways was encountered to a similar extent as on substrates patterned with the entire protein.

16.2 Surface functionalization The deposition of chemical cues to solid substrates by a mere physisorption step, e.g., by adsorption of polymers or proteins from solution, is the simplest way to generate a specific functionality. However, the interaction forces involved are typically weak and, hence, the coatings are prone to detachment, delamination, or desorption, e.g., adhesion proteins bound to surfaces via hydrophobic interactions can be displaced by other (small) hydrophobic or amphiphilic molecules. Even though the many weak Coulombic interactions between the charges along a polyelectrolyte molecule adsorbed to the countercharges at a correspondingly modified surface give a relatively stable layer, a gradual dissolution of the adsorbed polymer by the interference of small ions renders this coating inherently instable.

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The introduction of covalent bonds between groups on the substrate surface and correspondingly reactive moieties of the molecule to be attached provide the coating with the desired long-term stability, even under chemical or physical stress, e.g., against electrolyte attack, temperature gradients, or flow shear-stress conditions. For low molecular mass systems the `grafting-to' approach, i.e., the coupling of a molecule through a reactive endgroup to the surface has proven to be a very versatile strategy for the introduction of a broad range of functionalities or a general modification of surface properties, e.g., the generation of hydrophobic/ hydrophilic interfaces, the control of wettability, the deposition of surface lubrication layers. For chain- or rod-like molecules, the gain in free energy by bond-formation between their reactive head-group and the substrate can overcompensate the loss of entropy of the molecule in going from solution to the surface and, thus, allows for the formation of highly organized, even crystalline surface coatings. Firstly, as demonstrated for long alkyl-chains with multifunctional silane head groups (Sagiv, 1980), this concept of monolayer formation by a self-assembly process has gained a widespread popularity through the chemically more versatile thiol derivatives (Nuzzo and Allara, 1983). While the latter are strongly binding to (noble) metal surfaces and, thus, allow for the stable modification of, e.g., electrode surfaces, the former silyl-derivatives are well-suited to the coupling to oxide surfaces, i.e., for the modification of glasssubstrates or non-metallized oxide gate electrodes of field-effect transistors. If silane- (or thiol-)derivatives also exhibit a reactive end-group, further coupling steps can lead to more complex supramolecular interfacial architectures. An example of such a strategy is given in Fig. 16.3. The surface silanol(SiOH-)groups of a glassy substrate are coated with APTMS (aminopropyltrimethoxysilane) resulting in an amine-terminated surface layer. A so-called hetero-bifunctional cross-linker, i.e., a small molecule with one reactive group specifically designed to couple to amine groups (cf the structure formula in Fig. 16.3) can be attached next, leading to a reactive surface capable of binding ± via its second reactive group ± e.g., the SH-group of a cysteine-terminated peptide or protein. In the example given in Fig. 16.3, the peptide consists of the binding domain of the B2 chain of laminin. Thus, a carpet of peptide motifs is generated which has been shown to be highly attractive for the adhesion of randomly seeded hippocampal neurons (Matsuzawa et al., 1996). This concept offers a number of advantages in addition to being a simple recipe that involves only commercially available coupling reagents. The choice of the linker system can be optimized in terms of its chemical nature and its molecular architecture to the extent that chemical degradation of the final protein layer by a denaturing interaction with the surface can be minimized. Moreover, the aminosilane layer can be easily patterned (cf. below) by photolithographic strategies. By choosing the appropriate UV-wavelength, the Si-O-Si bond can be selectively hydrolyzed again regenerating free silanol groups in the exposed areas for the next self-

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16.3 Peptide immobilization on a surface using a heterobifunctional crosslinker. The silanol-groups of a glass substrate are modified with aminosilane groups, leaving an amine-terminated surface. These amine groups subsequently react with one of the functional groups of a crosslinker molecule, thus immobilizing it on the surface while the second functional group remains free. During the next step, the surface is incubated with a peptide or protein containing a terminal cystein residue. The SH group of the cysteine then forms a covalent bond with the free reactive group of the crosslinker molecule, leaving a carpet of peptide chains on the surface.

assembly step, e.g., with a silane derivative of another end group functionality (Dulcey et al., 1991). The `grafting-to' principle has been shown to work well for low molecular mass molecules. However, in cases where longer, polymeric systems were to be end-grafted, i.e., were to be coupled to the substrate via one of their two end groups, this concept failed for mainly two reasons (RuÈhe and Knoll, 2000). Firstly, the statistical distribution of the two end groups within the random coil of a polymer in solution makes it more and more unlikely that for an increased molecular mass polymer, that one of these will reach the surface and form a required covalent bond with a surface reactive moiety. Secondly, the increase in surface coverage by polymer chains that have already bound to the surface generates an increasing kinetic barrier to the next chain approaching the interface from solution. The osmotic pressure generated by the bound layer slows down the penetration of the endgroups of the new chains and thus prevents them from forming a stable bond with the surface. The `grafting-to' approach therefore leads to only rather thin surface coatings in the range of only a few nanometers.

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If dense polymer brushes are required, a novel concept originally proposed by DeGennes, and experimentally realized by the RuÈhe group can be used. Instead of coupling the whole polymer to the surface, only the small initiator molecule is covalently bound and thus leads to a functional surface coating of high density (Prucker and RuÈhe, 1993). The induction of the polymerization reaction by activating surface-initiators by heat or light in the presence of monomers can then lead to polymer brushes of very high graft densities (down to a chain-chain separation of d0 ˆ 2±3 nm) and with individual chains reaching a molecular mass of 106 Da. This concept allows the use of a wide variety of monomers, thus providing a recipe for the generation of very different surface cues. An additional advantage is the possibility of activating the initiator by photons, thus, enabling the generation of patterned surface coatings using a suitable mask for photo-initiation. This way, not only attractive coatings but also highly cell-repellent areas have been generated on solid surfaces, e.g., by the growth of a patterned brush of only 20 nm poly(styrene), the coated areas on the surface of a transistor chip were completely free from adherent cells after random seeding. Only the protected uncoated areas were covered by a dense monolayer of cells (Prucker et al., 1998).

16.3 Patterning of chemical surface cues 16.3.1 Photolithographic patterning Photolithographic techniques are well established for the mass production of silicon chips with a resolution and alignment precision of different surface areas in the sub-m range. Patterns are created by the partial exposure of photoresist surfaces to UV light through a chromium-quartz mask, followed by a chemical development step during which the exposed areas are etched away. Afterwards, the surface in these regions can be chemically modified, while the resist-covered sections are inaccessible. Subsequent removal of the photoresist with organic solvents leaves well-defined patterns of modified and unmodified surface areas (Fig. 16.4). Although the technique was initially developed for inorganic materials, Kleinfeld and coworkers successfully applied it to pattern thin films of organic molecules (Kleinfeld et al., 1988). One drawback to this system for application in cell patterning is that the solvents and development solutions are chemically relatively aggressive. It was therefore unclear whether the method can also be utilized to pattern biological molecules, which are rather sensitive with respect to the surrounding milieu. In 1993, Clark and coworkers (Clark et al., 1993) found an elegant solution to circumvent this problem by patterning the adhesion molecule laminin indirectly through a photolithographical method. In the initial patterning step, methyl groups were bound to defined surface areas while the remainder were left uncoated. As laminin adsorbs preferentially to hydrophobic surfaces, incubation of the substrates with a laminin solution resulted in a pattern of laminin adsorbed

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16.4 Scheme of the steps involved in surface patterning by photolithography. A quartz substrate onto which a thin layer of photoresist has been applied by spin-coating (a) is irradiated with UV light through a photomask (b). During the following development step, the exposed areas are etched away (c). The surface is then flooded with the desired molecules (e.g. alkyl-trichlorosilanes), which bind to both the photoresist-covered and uncovered areas (d). After removal of the photoresist with organic solvents (e), the surface is incubated with a second type of molecules (e.g. aminosilanes), for which only the previously covered surface areas are available for binding (f). Note that the figure is not drawn to scale, the photoresist layer is  1 m thick, while the alkanes and amines form layers of molecular thickness. Adapted from: Kleinfeld et al., Journal of Neuroscience 8, (1988), 4098±120.

to the methylated surface areas against uncoated quartz. The protein coated tracks were shown to guide the outgrowth of neurites from chick embryo brain neurons, indicating that at least the epitope(s) required for this function had retained their native conformation. Since then, standard photoresist techniques have been adapted to generate micropatterns of proteins on glass by using lift-off and plasma-etching techniques (Tai and Buettner, 1998; Sorribas et al., 2002).

16.3.2 Photochemical patterning Photochemical methods can be used to pattern self-assembled monolayers (SAM) or thin films of organic molecules by exposing the surface to UV light

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through a photomask or a metal mask. Illumination with UV light causes oxidation of the molecules in the exposed areas (for example alkanethiolate oxidize to alkanesulfonate) which alters their chemical properties, e.g., their solubility. Immersion of the patterned substrate in a solution containing another organic molecule results in the modification of the illuminated region by a second monolayer (Dulcey et al., 1991) while the non-illuminated region remains inert. An appropriate choice of molecules for the two layers allows for the creation of patterned surfaces with contrasting adhesive properties suitable for controlling cell adhesion to defined sites. The method has been utilized by Ravenscroft and coworkers who irradiated a thin surface bound film of the cytophilic molecule DETA (trimethoxysilylpropyldiethylenetriamine) through a photolithographic mask, thus creating reactive hydroxyl-groups in the illuminated areas. These reacted with the cytophobic silane 13F (tridecafluoro-1,1,2,2-tetrahydrooctyl-1-dimethylchlorosilane) in a second step which resulted in complementary anti-adhesive surface areas (Ravenscroft et al., 1998).

16.3.3 Plasma micropatterning Plasma polymerization is a method in which gaseous monomers, excited by a plasma, react and precipitate onto two- or three-dimensional substrates as a highly crosslinked polymer layer. Through the right choice of monomers, plasma power input and deposition time, it is possible to control a number of properties of the deposited layer, such as its hydrophilicity, chemical nature, density of functional groups and surface energy. This has been exploited to systematically render surfaces attractive or repulsive to cell attachment. Deposition of n-hexane films, which resulted in cell repulsion, through a transmission electron microscope mask, has been used for the creation of a polymer pattern on permissive tissue culture dishes. Groups in Sheffield and Aberdeen were able to demonstrate the utility of these surfaces by growing neuronal cells on these substrates and show that they adhered selectively to the protected areas of the culture dish (Bullett et al., 2001; Mitchell et al., 2002).

16.3.4 Microcontact printing Microcontact printing (CP) was initially developed by the Whitesides group who used it to print monolayers of alkanethiols onto gold substrates (Kumar and Whitesides, 1993). An elastomeric stamp is employed to create patterns of organic molecules on planar surfaces. The application of this technique to the investigation of cell-substrate interactions has mainly focused on endothelial cell adhesion and control of neuronal process outgrowth for the creation of defined neuronal networks (Singhvi et al., 1994; Branch et al., 1998, 2000; James et al., 1998; Scholl et al., 2000; Kam et al., 2001; Yeung et al., 2001).

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16.5 SEM image of a grid-shaped PDMS stamp for microcontact printing, kindly provided by Tanja Decker.

Similar to photochemical patterning methods, the procedure starts with a photolithography step to produce the mould (master stamp). Photoresists with a high aspect ratio and a thickness of more than 5 m are used to realize high relief topographical structures in the photoresist after illumination and development steps. Polydimethylsiloxane (PDMS), a viscous liquid that polymerizes to an elastomeric, rubber-like material, is poured into these molds and allowed to cure, thus solidifying to a stamp containing the topographical pattern complementary to that of the mold (Fig. 16.5). The PDMS stamp is then wet (`inked`) with a solution of organic molecules, dried and placed in contact with a surface. Similarly as with any macroscopic stamp used in everyday life, the organic molecules are transferred only at those regions where the stamp contacts the surface, leaving the pattern defined by the stamp (Fig. 16.6). As the procedure involves a drying step which may cause the denaturation of some proteins, an indirect two-step process was developed by the group of Gary Banker (Oliva et al., 2003). Bacterial protein A, which is known to have a relatively robust tertiary structure, was printed onto a tissue culture dish. Subsequently, the protein of choice ± the cell adhesion molecule L1 ± was allowed to bind to the surface from solution. Selective deposition on the pattern was achieved by coupling L1 to the Fc domain of an antibody molecule, which is known to bind protein A with high affinity. However, this two-step technique is not always necessary, as a range of matrix proteins has been shown to retain their native conformation during printing as indicated by their sustained ability to mediate cell attachment or synapse formation (Cornish et al., 2002). This fact is thought to be attributable to the formation of densely packed protein crystals during the slow drying step, which allows for the preservation of secondary and tertiary structures. The patterned surface can be in the range of several cm2 in size, while features can have an edge resolution in the sub-m range.

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16.6 Surface modification by microcontact printing. (a) Schematic representation of the steps involved: the stamp is wet with the inking solution of choice (1) and dried, leaving a thin layer of material on the stamp surface (2). Usually, a cell attractant material is used. The stamp is then pressed to the substrate surface (3), leaving the inking material only on the regions defined by the stamp topography (4). (b) SEM image of a surface onto which a grid pattern was printed using the stamp shown in Fig. 16.5. A blend of extracellular matrix protein was used as an inking solution and printed onto a polystyrene substrate.

16.4 Synaptic connections in patterned neuronal networks: communication along predefined pathways One of the most striking properties of the central nervous system lies in its ability to process and store information. Neurons communicate through electrical signals (action potentials) which originate at the cell body as a result

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of integrated signal input, and can be passed on to other cells through synaptic contacts. Synapses can be divided into two major classes, electrical and chemical synapses, which are distinct in structure and functionality. Electrical synapses, also called gap junctions, consist of a pore spanning the lipid bilayers of two adjacent cells. This pore has an internal diameter of about 2 nm, and is thus large enough for many ions and small metabolites to pass. Consequently, electrical currents can flow between the two connected cells, such that differences in the membrane potential of one cell have a direct effect on the potential of a coupled cell(s). Through this immediate electrical connection, signals are transduced directly with almost no time delay. Therefore, electrical synapses play a central role in the transmission of signals important for actions that need to be rapid, like flight reflexes. Additionally, electrical connectivity has been implicated in the synchronization of oscillatory activity in the brain (Galarreta and Hestrin, 1998). Chemical synapses are the second major type, and typically form between the axon of the presynaptic cell and a dendrite of a postsynaptic cell. They do not allow for the direct flow of current as the pre- and postsynapse are separated by a gap of about 30 nm, the synaptic cleft. If an action potential traveling down the axon of the presynaptic cell reaches the presynaptic terminal, transmitter molecules are released into the synaptic cleft. These can bind to receptors in the postsynaptic membrane, which directly or indirectly leads to the opening of an ion channel. The resulting ion flux alters the transmembrane potential and facilitates or suppresses the generation of action potentials in the postsynaptic cell (Nicholls et al., 1992). Signal transmission at a chemical synapse is unidirectional from pre- to postsynapse and occurs with a time delay of several ms. This delay arises due to the time it takes for the release of the synaptic vesicles from their anchorage in the cytoskeleton and their fusion with the presynaptic membrane. While signal transmission through electrical synapses is thought to be relatively constant, chemical synapses are subject to a wide range of modulatory mechanisms that may strengthen or weaken signal transmission in response to previous input patterns. Such activity-dependent modifications of synaptic efficacy are central on a number of capabilities of the CNS: While long-term changes are believed to be involved in learning and the formation of memory (Bliss and Lomo, 1973; Fitzsimonds et al., 1997; Ganguly et al., 2000), shortterm plasticity has been implicated in the processing of incoming signals, e.g., recognition of different input patterns or contrast adaptation (Chance et al., 1998; Kaplan et al., 2003). Sensitivity to particular patterns of sensory input is additionally attributed to features of the microcircuitry through which neurons, within a tissue, are interconnected. Different afferent signals may be attenuated or amplified depending on the `wiring' of the network, e.g., the extent to which feedback loops or reciprocal connections prevail (Muller et al., 1997; Misgeld et al., 1998; Nelson and Turrigiano, 1998; Manor et al., 1999).

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The investigation of synaptic plasticity in vitro is mostly performed by patchclamp recordings between two synaptically connected neurons in brain slices. However, the extreme complexity of neuronal connectivity within this organ renders the investigation of single contacts extremely difficult as the synaptic input from outside a constellation of interest interferes with experimentally applied signals and impedes reproducibility. Low-density neuronal cultures, in which isolated islands of two or three synaptically connected cells are found, have yielded valuable insights in synaptic modulation at the level of single synaptic contacts (Bi and Poo, 1999; Ganguly et al., 2000). Such cultures, however, do not allow for the experimental determination or alteration of the connectivity pattern, such that the impact of network architecture on signal transduction and processing cannot be studied systematically. Addressing these issues, the approach of patterned neuronal cell culture emerged. In such cultures, it is possible to grow neuronal networks of extremely low complexity, because only a limited number of pathways between the adherent cells are open for the formation of synaptic connections. Moreover, the approach also allows experimental design of the network architecture, such that linear arrangements, feedback circuits or branching pathways can be studied (Fig. 16.7). One problem with the application of patterned neuronal cultures as a model

16.7 Different possible network geometries that may be realized by growth on patterned substrates to study the impact of network architecture on signal processing. Adapted from Steward, Nature 427, 601±604 (2004).

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system lies in the fact that for a long time cells were found either to comply well with the respective surface geometry and show physiological deficits (Lauer et al., 2002; Matsuzawa et al., 2000; Stenger et al., 1998) or to develop normally while overgrowing the pattern boundaries (Ma et al., 1998; Liu et al., 2000). In particular, the formation and maturation of chemical synapses in culture fully compliant with a micropattern proved to be a challenge (Ravenscroft et al., 1998; OffenhaÈusser et al., 1997). Potential reasons for these problems may lie in the synthetic groups frequently used for surface modification which do not mediate attachment through cell adhesion molecules, but instead through other forces, e.g., electrostatic interactions. Alternatively, some of the chemical groups may have mild cytotoxic effects which impair cellular development. It has also been speculated that the geometry itself may be a factor modulating neuronal development, and that limitation of neurite outgrowth to a four-way-grid pattern is a state impeding neuronal maturation (Ravenscroft et al., 1998). Consistent with this hypothesis, Wyart and collegues reported that patterns of polylysine severely impaired cellular physiology if the adhesion sites offered for neuronal attachment were smaller than 80 m. This is an area significantly larger than a typical cell body and consequently allows the extensive outgrowth and branching of neuritis (Wyart et al., 2002). However, it was recently shown that cells highly compliant to grid patterns displayed normal physiological characteristics similar to cells grown on homogeneous control substrates. Moreover, cells were found to connect through electrical and chemical synapses that developed normally by all the criteria tested, such that communication along the predefined pattern could be observed. The pattern had been created by applying a blend of extracellular matrix proteins to a background of polystyrene by microcontact printing, such that a relatively natural substrate was offered for neurite outgrowth (Vogt et al., 2003). Moreover, the encountered synapses exhibited several forms of short-term plasticity similar to that encountered in the intact brain (Vogt et al., 2005). Short-term plasticity, the modification of synaptic strength for seconds to minutes, in response to different stimulus patterns, has been shown to be essential for a number of abilities of the mammalian brain such as the recognition of different input patterns. Encountering these features in the system while cell attachment and connectivity are highly restricted to an underlying pattern qualifies it as a suitable model system: Features basic to neuronal signal processing are preserved while the network geometry can be experimentally varied, allowing the investigation of the impact of different types of circuitry on the way incoming signals influence the basal network activity.

16.5 Conclusion Cellular adhesion to a surface can be experimentally manipulated through either chemical or topographical signals. The use of patterned surfaces consisting of

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small adhesive areas against a cell-repellent background allows the precise positioning of cell bodies thereon. In neuronal cultures, the outgrowth of neurites and thus the connectivity pattern of a neuronal network can additionally be guided by adhesive pathways interconnecting these adhesion sites. Such networks of defined geometry provide a basis for a number of biotechnological applications as well as highly defined experiments on cellular interactions, network behavior and neurocomputing. First steps towards functional networks of highly defined geometry have been taken. Highly compliant neuronal networks can be realized on patterned surfaces in which communication through mature synaptic contacts can be observed along the pattern pathways. A further challenge will be the demonstration of long-term modifications of synaptic strength in patterned networks, which are thought to underlie processes like learning and memory formation. A system reproducing these abilities of the central nervous system while network geometry connectivity pathways are experimentally controlable will be a valuable tool for the investigation of neuronal signal processing.

16.6 References Bi, G, Poo, M, Nature 401, (1999), 792±6. Bliss, T, Lomo, T, Journal of Physiology and Behavior 232, (1973), 331±356. Bohanon, T, Elender, G, Knoll, W, Koberle, P, Lee, JS, OffenhaÈusser, A, Ringsdorf, H, Sackmann, E, Simon, J, Tovar, G, Winnik, FM, J Biomater Sci Polym Ed. 8, (1996), 19±39. Branch, DW, Corey, JM, Weyhenmeyer, JA, Brewer, GJ, Wheeler, BC, Medical & Biological Engineering & Computing 36, (1998), 135±41. Branch, DW, Wheeler, BC, Brewer, GJ, Leckband, DE, IEEE Transactions on Biomedical Engineering 47, (2000), 290±300. Britland, S, Morgan, H, Wojiak-Stodart, B, Riehle, M, Curtis, A, Wilkinson, C, Experimental Cell Research 228, (1996), 313±25. Bullett, N, Short, RD, O'Leary, T, Beck, AJ, Douglas, CWI, Cambray-Deakin, M, Fletcher, IW, Roberts, A, Blomfield, C, Surface and Interface Analysis 31, (2001), 1074±76. Carter, S, Nature 208, (1965), 1183±7. Chance, F, Nelson, SB, Abbott, LF, J. Neurosci. 18, (1998), 4785±99. Clark, P, Connolly, P, Curtis, AS, Dow, JA, Wilkinson, CD, Development 108, (1990), 635±44. Clark, P, Britland, S, Connolly, P, J Cell Sci 105 (Pt 1), 203±12. Corey, JM, Wheeler, BC, Brewer, GJ, Journal of Neuroscience Research 30, (1991), 300±7. Cornish, T, Branch, DW, Wheeler, BC, Campanelli, JT, Molecular and Cellular Neurosciences 20, (2002), 140±53. Curtis, A, Wilkinson, C, Biomaterials 18, (1997), 1573±83. Curtis, A, Wilkinson, CD, J Biomater Sci Polym Ed. 9, (1998), 1313±29. Dulcey, C, Georger, JH Jr, Krauthamer, V, Stenger, DA, Fare, TL, Calvert, JM, Science 252, (1991), 551±4.

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Fitzsimonds, RM, Song, HJ and Poo, MM, Nature 388, (1997), 439±48. Fromherz, P, Schaden, H, Vetter, T, Neuroscience Letters 129, (1991), 77±80. Fromherz, P, Schaden, H, European Journal of Neuroscience 6, (1994), 1500±4. Galarreta, M, Hestrin, S, Nature Neuroscience 1, (1998), 587±94. Ganguly, K, Kiss, L, Poo, M, Nat. Neurosci. 3, (2000), 1018±26. Graf, J, Iwamoto, Y, Sasaki, M, Martin, GR, Kleinman, HK, Robey, FA, Yamada, Y, Cell 48, (1987), 989±96. Hammarback, J, Palm, SL, Furcht, LT, Letourneau, PC, Journal of Neuroscience 13, (1985), 213±20. James, C, Davis, RC, Kam, L, Craighead, HG, Isaacson, M, Turner, JN, Shain, W, Langmuir 14, (1998), 741±4. Jung, D, Kapur, R, Adams, T, Giuliano, KA, Mrksich, M, Craighead, HG, Taylor, DL, Critical Reviews in Biotechnology 21, (2001), 111±54. Kam, L, Shain, W, Turner, JN, Bizios, R, Biomaterials 22, (2001), 1049±54. Kaplan, M, Wilcox, KS, Dichter, MA, Synapse 50, (2003), 41±52. Kleinfeld, D, Kahler, K, Hockberger, P, Journal of Neuroscience 8, (1988), 4098±120. Kleinman, H, Luckenbill-Edds, L, Cannon, FW, Sephel, GC, Analytical Biochemistry 166, (1987), 1±13. Kleinman, H, Sephel, GC, Tashiro, K, Weeks, BS, Burrous, BA, Adler, SH, Yamada, Y, Martin, GR, ANN. N.Y. Acad. Sci. 580, (1990), 302±10. Kumar, A, Whitesides, GM, Applied Physics Letters 63, (1993), 2002±4. Lauer, L, Vogt, A, Yeung, C, Knoll, W, Offenhausser, A, Biomaterials 23, (2002), 3123±30. Letourneau, P, Developmental Biology 44, (1975), 92±101. Liu, Q, Coulombe, M, Dumm, J, Shaffer, K, Schaffner, A, Barker, J, Pancrazio, J, Stenger, D, Ma, W, Developmental Brain Research (2000), 223±31. Ma, W, Liu, QY, Jung, D, Manos, P, Pancrazio, JJ, Schaffner, AE, Barker, JL, Stenger, DA, Brain Research. Developmental Brain Research 111, (1998), 231±43. Manor, Y, Nadim, F, Epstein, S, Ritt, J, Marder, E, Kopell, N, Journal of Neuroscience 19, (1999), 2765±79. Matsuzawa, M, Liesi, W, Knoll, W, Journal of Neuroscience Methods 69, (1996), 189±96. Matsuzawa, M, Tabata, T, Knoll, W, Kano, M, Eur J Neurosci 12, (2000), 903±10. Merz, M, Fromherz, P, Adv. Mater. 2, (2002), 141±4. Misgeld, U, Zeilhofer, HU, Swandulla, D, Cellular and Molecular Neurobiology 18, (1998), 29±43. Mitchell, S, Emmison, N, Shard, AG, Surface and Interface Analysis 33, (2002), 742±7. Muller, T, Swandulla, D, Zeilhofer, HU, Journal of Neurophysiology 77, (1997), 3218±25. Nelson, S, Turrigiano, GG, Nature Neuroscience 1, (1998), 539±41. Nicholls, J, Martin, AR, Wallace, BG, From Neuron to Brain S Associates, Sunderland, Mass. (1992). Nuzzo, R, Allara, DL, Journal of the American Chemical Society 105, (1983), 4481. OffenhaÈusser, A, Sprossler, C, Matsuzawa, M, Knoll, W, Neuroscience Letters 223, (1997), 9±12. Oliva, AJ, James, CD, Kingman, CE, Craighead, HG, Banker, GA, Neurochemical Research 28, (2003), 1639±48. Prucker, O, RuÈhe, J, Materials Research Society Symposium Proceedings 304, (1993), 1675. Prucker, O, Schimmel, M, Tovar, G, Knoll, W, RuÈhe, J, Advanced Materials 10, (1998), 1073±7.

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Ranieri, J, Bellamkonda, R, Bekos, EJ, Gardella, JA Jr, Mathieu, HJ, Ruiz, L, Aebischer, P, Int J Dev Neurosci 12, (1994), 725±35. Ravenscroft, MS, Bateman, KE, Shaffer, KM, Schessler, HM, Jung, DR, Schneider, TW, Montgomery, CB, Custer, TL, Schaffner, AE, Liu, QY, Li, YX, Barker, JL, Hickman, JJ, Journal of the American Chemical Society 120, (1998), 12169±77. Rosenberg, M, Science 139, (1963), 411±12. RuÈhe, J, Knoll, W, Supramolecular Polymers A Ciferri, New York (2000) 565±613. Ruoslahti, E, Pierschbacher, MD, Cell 44, (1986), 517±18. Ruoslahti, E, Pierschbacher, MD, Science 238, (1987), 491±7. Sagiv, J, Journal of the American Chemical Society 102, (1980), 92. Scholl, M, Sprossler, C, Denyer, M, Krause, M, Nakajima, K, Maelicke, A, Knoll, W, OffenhaÈusser, A, Journal of Neuroscience Methods 104, (2000), 65±75. Singhvi, R, Kumar, A, Lopez, GP, Stephanopoulos, GN, Wang, DI, Whitesides, GM, Ingber, DE, Science 264, (1994), 696±8. Sorribas, H, Padeste, C, Tiefenauer, L, Biomaterials 23, (2002), 893±900. Stenger, DA, Hickman, JJ, Bateman, KE, Ravenscroft, MS, Ma, W, Pancrazio, JJ, Shaffer, K, Schaffner, AE, Cribbs, DH, Cotman, CW, Journal of Neuroscience Methods 82, (1998), 167±73. Tai, H, Buettner, HM, Biotechnology Progress 14, (1998), 364±70. Valentini, R, Vargo, TG, Gardella, JA Jr, Aebischer, P, J Biomater Sci Polym Ed 5, (1993), 13±36. Vogt, AK, Lauer, L, Knoll, W, OffenhaÈusser, A, Biotechnology Progress 19, (2003), 1562±8. Vogt, AK, Wrobel, G, Meyer, W, Knoll, W, OffenhaÈusser, A, Biomaterials 23, (2005), 3123±30. Wyart, C, Ybert, C, Bourdieu, L, Herr, C, Prinz, C, Chatenay, D, Journal of Neuroscience Methods 117, (2002), 123±31. Yeung, CK, Lauer, L, OffenhaÈusser, A, Knoll, W, Neuroscience Letters 301, (2001), 147±50.

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Controlled cell deposition techniques

C M A S O N , University College London, UK

17.1 Introduction Important challenges for the successful engineering of human tissues and organs include elucidating how the cellular microenvironment interacts with cell function ± a bidirectional, dynamic and highly complex set of interactive relationships. Thus the ability to be able to position, manipulate, orientate the architecture and control multiple cell types together with their extracellular matrix at the cellular and molecular level of resolution is paramount for the fabrication of true-to-life three-dimensional living constructs. This chapter focuses on three key areas: 1. 2. 3.

In-vivo and in-vitro cell interactions. Two-dimensional in-vitro controlled cell deposition. Three-dimensional in-vitro controlled cell deposition.

A multidisciplinary approach to the topic is mandatory since only by understanding the in-vivo embryology, histology, biochemistry and physiology together with a knowledge of cell-substrate interaction, is it possible to begin to evolve the controlled cell deposition techniques required to fabricate accurately representative two-dimensional and three-dimensional de novo cellular structures. Two-dimension controlled cell deposition has its origins in traditional cell culture where cell monolayers are seeded onto flat surfaces in sterile flasks. The history of animal cell culture dates back approximately one hundred years to the pioneering work by Harrison and Carrel (Harrison, 1907; Carrel, 1912). Over the past few decades, a paradigm shift has occurred due to the convergence of cell culture with electronic semiconductor technology allowing cells to be accurately patterned onto substrates. Three-dimensional controlled cell deposition techniques are altogether much more recent and as such are at an extremely early stage of their development. However, the advancing fields of regenerative medicine and, in particular, tissue engineering are highlighting the potential demand for such technology.

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17.2 In-vivo and in-vitro cell interactions The complex mechanisms whereby cells interact and function as tissues and organs has evolved over hundreds of millions of years. Understanding these processes in vivo in conjunction with how the same processes interact with synthetic polymers is key to controlling cell deposition in vitro.

17.2.1 Animal cells and tissues The majority of cells in the body are not isolated individual entities but are grouped together to form tissue and organs. Histologists classically classified most cells into two discrete groups; epithelial and mesenchymal cells depending upon their location and orientation. Epithelial cells principally cover surfaces (e.g. skin epithelium) and line cavities and hollow tubes (e.g. endothelium lining blood vessels). A feature common to all epithelial cells is that they form continuous sheets or layers with their outer cell membranes tightly butted up against one another like paving slabs forming a well-made pavement. In stark contrast, the mesenchymal cells (or connective tissue cells) which make up the rest of the body are widely spaced apart in a ground substance with the generic name extracellular matrix (ECM). Whilst cells are the essential unit of living material, it is the ECM which is largely responsible for giving individual tissues and organs their unique and specific properties. Thus in vivo, cells interact both with one another and with their surrounding ECM. It is therefore important that cells grown in vitro have the same orientation, density and interactions with their environment if they are to be truly representative of their in-vivo counterparts. Cells have traditionally been cultured as monolayers of single cell types which in general are not totally representative of the same cells living in a three-dimensional tissue or organ. This arrangement, however, has the advantage of simplicity and can `give important information on the characters acquired by tissues liberated from the control of the organism from which they were derived' (Carrel, 1912), e.g., physiological processes including signal transduction and the regulation of gene expression. The alternative to single cell type monolayers is to fabricate whole tissues and organs in vitro, in which collections of different cell types with a temporary artificial ECM (scaffold) are cultured together in a single bioreactor. This is the approach adopted by today's tissue engineers (Stock and Vacanti, 2001).

17.2.2 Extracellular matrix The ECM is the non-cellular structural material that surrounds mesenchymal cells in order to form tissues and organs. There are two distinct varieties of ECM. The first variety fully encapsulates mesenchymal cells, the second type is

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in the form of a thin amorphous lamina which creates the interface between the basal surface of epithelial cells and their directly underlying mesenchymal tissue (Bard, 1990). Produced by the mesenchymal cells, the ECM has a number of roles including structural support and provision of environmental signals to control site-specific cellular regulation (Junqueira and Carniero, 2003). CellECM interactions are bidirectional and dynamic in nature. Thus, mesenchymal cells directly influence the composition of their surrounding matrix whilst the matrix feeds back to the encapsulated cells. This feedback can be either biochemical or biomechanical in nature. Thus a basic understanding of cellECM interaction is vital if researchers are to be able to accurately mimic in-vivo situations in vitro. In addition, in cell culture, there is an added complexity, namely, cell-synthetic polymer interaction. The majority of animal cells are only capable of proliferating when attached to a surface, aptly named anchorage dependent cells. Therefore, in-vitro studies typically involve using a polymer surface (e.g. polystyrene) upon which the cells can anchor. To aid cell attachment, manufacturers of cell culture plastic-ware pre-treat the surfaces, since only the uppermost layers of the polymer are in direct physicochemical contact with the biological environment (Voger, 1993). In general, the adhesion of animal cells to polymer surfaces occurs in two distinct steps. Firstly, the cell on approaching a surface expands out a temporary pseudopodial extension in order to examine the surface for suitable protein ligands prior to forming an initial temporary focal attachment. This is followed by the cell spreading itself out over the surface and creating more permanent local attachments (Mosher, 1993). The adhesion process is created by a combination of non-specific and specific interactions (Voger, 1993). The nonspecific interactions include hydrogen bonding and van der Waal's forces (Parsegian, 1981), whilst specific adhesion involves receptor-ligand bonding, e.g., integrins, a family of proteins that bind to ECM molecules such as fibronectin, collagen and laminin (Irvine et al., 2002).

17.2.3 Solid-fluid interface As a generalisation, due to their tertiary structure, mammalian proteins prefer to be in an aqueous environment. However, when an aqueous solution of proteins is exposed to a solid surface, the proteins have a tendency to spontaneously accumulate (adsorb) at the solid-fluid interface due to a combination of their degree of hydrophobicity, electrical charge and polar forces. This propensity to accumulate following collision with the interface is dependent upon a number of variables, including the nature of the solid surface, temperature, aqueous solution characteristics (e.g. protein concentration and flow) and the duration of exposure (Norde, 1995; Norde and Haynes, 1996). The duration of exposure required to trigger adsorption to a synthetic surface is in the order of seconds (Vroman, 1987). Adsorption of proteins to a synthetic surface initiates the

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formation of a biopolymer film which in turn prepares the surface for cell attachment. Cell attachment quickly ensues in a matter of a few minutes. Adsorption is, however, not a simple process whereby once the synthetic surface is coated with protein the priming of the surface is totally completed. Instead, it is a dynamic process e.g. adsorbed proteins are replaced by other proteins in the solution and/or adsorbed proteins undergo conformational changes. For example, in vivo, adhesion of proteins to artificial surfaces is dominated by the Vroman effect in which the more abundant proteins are adsorbed first and then are replaced by less abundant proteins which have a higher affinity for the particular artificial surface (Vroman, 1974; Vroman and Adams, 1986). For example, when plasma is exposed to a glass surface, albumin (the most abundant protein) is adsorbed first, which is then replaced by fibrinogen (less abundant) which is in turn replaced by kininogen (least prevalent protein) (Pfeiffer et al., 1998; Jung et al., 2003). Finally, following adsorption, protein-adsorption kinetics may favour either conformational change and/or denaturing of the protein (Brash and Horbett, 1995). Understanding the cell-substrate interaction lies at the core of in-vitro controlled cell deposition, since cells in vivo are not simply stuck together with ECM `glue' in a random arrangement but are organised into diverse and distinctive patterns (Gumbiner, 1996). A detailed review of this topic has been undertaken by Saltzman (2000).

17.3 Two-dimensional controlled cell deposition techniques 17.3.1 Photolithography Photolithography creates a topographic pattern on a surface by lithography (the process of printing from a surface on which the printing areas are not elevated but are ink receptive) combined with an etching process. This approach has its roots firmly embedded in the semiconductor industry (Britland et al., 1992). Typically, the technique involves a suitable surface (e.g. borosilicate glass, fused quartz or polished silicon wafer) being spun coated with a thin photosensitive organic resist film (e.g. diazo-naphtho-quinone with a novolak base resin) (Nicolau et al., 1999). Using an exposure aligner, an opaque mask (typically a chrome film) of a desired final pattern is placed over the photosensitive resist film. Ultraviolet radiation is then directed at the unmasked areas resulting in the exposed areas of the film increasing in solubility upon the addition of a developer solution. Thus selective dissolution (wet etching) of the membrane can be achieved resulting in a photoresist pattern of membrane identical to the mask. Hydrofluoric acid is now deployed, since it can react only with the areas of exposed glass, a precise pattern is etched in the surface. The depth of the channel produced is proportional to duration of exposure. Etching of glass is isotropic, i.e., etching proceeds in all directions resulting in the undercutting of

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the mask. This results in the etched channel being `canoe' shaped in crosssection. The approximate channel width is, therefore, the width of the mask plus twice the etch depth. In contrast, if silicon is wet etched, the process is anisotropic due to the different reactivity of the different crystal planes. Overall the entire process of fabricating a chip using wet etching involves approximately 100 distinct steps (Greenwood, 2004). Whilst wet etching is the predominant etching technique there are, however, a host of alternatives, including laser ablation, micromachining, powder blasting, embossing, LIGA (Lithografie, Galvanoformung, Abformung ± lithography, electroplating, injection moulding) and plasma ion etching. The choice of etching technique depends upon a range of factors including the material to be etched, the desired channel shape, required pattern resolution and cost. Each of the methods has its advantages and disadvantages. For example, laser ablation vaporises the material and hence no undercutting of the mask, resulting in extremely precise square channels. The trade-off is that the resulting channel surface has a rough texture. Likewise, powder blasting, an anisotropic process, also results in a rough surface. Micromachining is performed by using a computer controlled hard milling wheel to erode the channels, unfortunately the size of the milling wheel limits the degree of resolution (<100 microns is difficult to achieve). Finally, if required, a combination of the above techniques can be employed in combination to produce a complex or deep pattern of channels (Greenwood, 2004). After the chosen method of etching has been completed the patterned substrate can either be immediately deployed or if desired, be bonded to another layer of the same material to produce a microfluidic chip. This requires the mating surfaces to be highly polished prior to being accurately placed together and fused in a kiln (approximately 600 ëC for borosilicate or 4000 ëC for quartz). If increased channel depth is required two matched etched surfaces can be fused face to face to produce double height channels. The chips are now ready for final preparation to make the channels `cell-friendly'. For example, coating the channel surfaces with cell adhesion proteins, which results in a substrate accurately patterned for selective cell deposition. Upon incubation with living cells suspended in culture medium, the cells attach to the surface only where the adhesion molecules are present, resulting in a monolayer of cells arranged in the desired pattern. Sophisticated versions of the basic technique have been used to produce accurate cell patterns on surfaces allowing in-vitro studies of cellsurface interactions and cell migration. For example, Lom, using a combination of photolithography, silane-coupling and protein adsorption, has fabricated patterned coverslips with amines, alkanes, and proteins with micrometre spatial resolution. Using patterned coverslips fabricated with different proteins then seeding neuroblastoma cells and utilising the transparent nature of the glass coverslips, direct visualisation of the cells was undertaken. Lom was able to demonstrate that cell attachment and neurite outgrowth displayed the following

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preferences: laminin, fibronectin, or collagen type IV > amine or glass > alkane or bovine serum albumin (Lom et al., 1993). Due to the origin of photolithography in the highly successful semiconductor industry, it has many advantages including being able to pattern at a submicron scale, ease of automation and scale-up, plus low individual item cost and furthermore, its research and development is underpinned by billion dollar commercial investment programmes. Another advantage is the relative ease with which chips can be cleaned for re-use, essential given their current high cost to fabricate. Various methods are available to rid the chips' surfaces of adsorbed proteins and cells without damaging the surface of the microchannels. These methods include flowing either weak hydrofluoric acid or sodium hydroxide solution through the channels. Blocking of channels, e.g., with dust, cell aggregates or precipitates is also an important problem, but again one that is resolvable using, for example, piranha solution (3:1 ratio of sulphuric acid and hydrogen peroxide). Sterilisation, an essential prerequisite for cell culture work, can be achieved by autoclaving. Provided chips are carefully maintained, their lifetime is in the order of one to two thousand experiments. There are, however, serious drawbacks when using the technique for life science research applications especially where volumes of units are presently minuscule when compared with the market for mass-produced silicon microchips. The major drawback for life science research is that current photolithographic fabrication requires cleanroom facilities and expensive manufacturing equipment. Furthermore, the standard reagents developed for microchip fabrication are both toxic to cells, and often hostile to cell attachment proteins, since biocompatibility is not an issue in semiconductor design. Despite the disadvantages, there are a number of proponents of the technology who have succeeded in producing micropatterns of cultured cells including Matsuda et al. (1990, 1995) with bovine endothelial cells, Nicolau et al. (1999) with avian neuronal and glial cells and Goessl et al. (2001a,b) with human polymorphonuclear leukocytes and vascular smooth muscle cells (Goessl et al., 2001c). Because of the disadvantages of using photolithography for life science applications a number of alternative technologies have been invented and developed that are both more cost effective and cell friendly.

17.3.2 Soft lithography The major alternatives to photolithography are based on a technique termed soft lithography, so called because a soft elastomeric material is used as a stamp or mould in order to transfer a pattern to a surface. However, soft lithography still requires the fabrication of a mask in order to produce a master mould or form. This is ideally produced using photolithograpy, but lower-cost alternatives have been described (Qin et al., 1996). The master mould or form is used to cast the

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soft elastomeric stamps, i.e., replica moulding (Xia et al., 1996). Whilst the photolithographic mask is not required to be made of chrome (i.e. a thin chromium film which is the most expensive step in conventional lithography), this step still adds cost and inconvenience and therefore acts as a barrier to entry for researchers to the overall process, estimated in 2001, to be in the order of $500 per square inch for features between 1±5 microns (Whitesides et al., 2001). Whilst photolithographically produced masks are optimal, researchers have devised alternative mask production techniques. For example, for the fabrication of masters with patterns requiring greater than 20 micron resolution, the mask can be manufactured using an alternative lower-cost and far more widely available technology, namely high resolution commercial laser printing. The pattern can be designed using a basic computer aided design (CAD) package and then printed onto an overhead transparency film using a commercial laserassisted image setting system. This alternative process takes only a few hours as opposed to about a week to obtain a conventional chrome mask produced by a commercial custom fabricator. Another advantage is that the mask is flexible and can therefore be used to pattern non-planar surfaces. Lastly, being very thin, the film masks can be stacked one on top of another in order to create new patterns. A drawback to this alternative mask technology is that the smallest feature that can be generated is limited by the resolving power of the image setting system which is significantly less than can be achieved using semiconductor industry-type photolithography (Qin et al., 1996). Today, chrome masks can provide submicron resolution whilst other alternative techniques tend to have resolutions of the order of 1±10 microns. These techniques, including a number of different alternatives by George Whitesides' group at Harvard University (Cambridge, MA, USA), allow masks to be fabricated over a variety of different pattern resolutions, e.g., using a commercial low power laser to ablate a sub 10 micron resolution pattern in a thin poly(methyl methacrylate) film doped with a dye (rhodamine B base) (Grzybowski et al., 1998). These alternative mask production techniques greatly reduce the barrier to entry and cost to such a point that the efficient evolution of patterns in a rapid prototyping manner is highly feasible, e.g., the design and fabrication of an elastomeric polymer form in less than 24 hours (Duffy et al., 1998). After the photomask fabrication step, the remainder of the soft lithography technique is relatively low cost since neither clean-room facilities nor expensive manufacturing plant are required. There are several excellent reviews on the topic of soft lithography and cell patterning, including articles by Kane et al. (1999), Takayama et al. (2000), Jung et al. (2001) and Park and Shuler (2003). The Jung review also contains a comprehensive history of cell patterning techniques from its origins in the 1960s. The major advantages of soft lithography emanate from the fact that unlike photolithography, the technique has been developed as a life science research

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application and hence the development of relatively low cost, easy to handle protein and cell friendly reagents. Other major benefits over photolithography are that because the final stamps are flexible, a variety of non-planar substrates can have patterns applied including submicrometre features on curved substrates with radii of curvature as small as 25 microns (photolithography only works on planar surfaces) (Jackman et al., 1995) also these alternative concepts for soft lithography microfabrication offer strategies for making three-dimensional microstructures with highly complex topologies (Jackman et al., 1998). After the production of the patterned master mould or form which will act as a template, a soft stamp or mould can be produced by casting a liquid polymer of the desired elastomer. The most commonly used polymers are siloxane based elastomers and one in particular is PDMS (polydimethylsiloxane) (Schmid and Michel, 2000). The advantages of using PDMS include excellent biocompatibility, ease of mass transfer of oxygen and carbon dioxide (non-polar gas permeability), flexibility (i.e. deformability to non-planar surfaces), optical transparency (greatly facilitates optical monitoring), durability (potential to reuse a stamp/mould many times) and a growing collection of scientific papers allowing both the process, and PDMS in particular, to be better characterised. A popular choice of PDMS for soft lithography is the Sylgard 184 silicone elastomer kit manufactured by Dow Corning (Midland, MI, USA). Using this kit, a PDMS construct can be produced by thoroughly mixing the base and curing agent in the ratio 9:1, degassing the homogeneous mixture prior to pouring it into a master mould and oven curing. Varying process parameters, e.g., temperature and setting time, result in constructs with controllable physical properties such as the degree of flexibility. PDMS is extremely biocompatible allowing its use in both in-vitro and in-vivo applications, indeed it already has a long history of use in the field of surgical implants such as breast implants and other plastic surgery procedures requiring tissue augmentation (Brandon et al., 2001; Kheir et al., 1998). A big advantage of PDMS is that once a structure is fabricated, its surface properties can be additionally modified, if required. For example, the surface of PDMS is intrinsically both hydrophobic and allows non-specific protein adhesion but can be modified. Possible modifications include making the surface hydrophilic by oxygen-plasma oxidisation (Chaudhury and Whitesides, 1991) or bonding functional groups to the surface capable of supporting the self-assembly of proteins, antibodies and mammalian cells (Lahann et al., 2003). The advantage of bonding functional groups to the surface overcomes a constant problem with modified PDMS surfaces, in that they tend to recover their original hydrophobicity (Linder et al., 2001). This is due to polymer interface surfaces being non-homogeneous, i.e., a surface can be considered as being composed of chemically distinct strata that blend into a continuum with the bulk of the polymer. These strata are not always stable in position and molecular configuration, hence exchanges of interface molecules with the bulk polymer occur (Voger, 1993).

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There are four main approaches to employing elastomeric moulds or stamps for the creation of cell patterns on a substrate: microcontact printing, microchannel patterning, laminar flow patterning and stencil patterning. Microcontact printing Microcontact printing of a pattern is based on physically transferring the preformed pattern from a soft elastomeric stamp onto a specially prepared surface (e.g. gold or silver). An excellent example of this technique is shown in Fig. 17.1. Typically the `ink' stamped onto the surface is a substance capable of rapidly reacting to form a self-assembling monolayer e.g. alkanethiolates of gold. This method of producing a pattern as a self-assembled monolayer (SAM) was developed by George Whitesides' group in the early 1990s (Kumar and Whitesides, 1993; Singhvi et al., 1994). Typically, the process involves the fabrication of a non-deformable master stamp using ideally photolithography to attain a high-resolution mask pattern. Initially, the surface is coated with a thin layer of photoresist, the photomask placed between an ultraviolet light source and the photoresist layer, ultraviolet exposure and finally rinsing to remove the now readily soluble exposed area of photoresist. PDMS is cast against the master stamp to create a relief pattern and allowed to cure. Being easily deformable the cured PDMS stamp can be separated with ease from the rigidly constructed master stamp. The master stamp is now ready to be re-used. In order to create a SAM pattern, the soft stamp is `inked' with a solution of alkanethiol in ethanol which is then allowed to dry. Interestingly, in Kumar's original paper, an actual ink pad was prepared by moistening a piece of lint-free paper with a solution of thiol in ethanol. Kumar notes that it was important to keep the ink pad as free of lint as

17.1 Micropatterned delayed brain tumour (DBT) cells. This image was captured 25 hours after the DBT cells were deposited onto a surface which was fabricated by first stamping 16-mercaptohexadecanoic acid onto a transparent gold film and then immersing the stamped film into a solution of ethylene glycol containing alkanethiol. Finally the SAMs were coated with fibronectin prior to plating with DBT cells. Line width dimensions from the left: first 2 lines, 20 microns; next 10 lines, 40 microns; next 10 lines, 80 microns; next 10 lines, 140 microns; last 3 lines, 10 microns. (Image contributed by Dr Elizabeth E. Endler Ph.D.)

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possible (Kumar and Whitesides, 1993). The stamp is then brought into conformal contact (i.e. maintaining the stamp's three-dimensional shape with no distortion due to pressure effects) with a gold substrate for a duration of a few seconds to a few minutes depending upon the exact reagents and pattern. The alkanethiol is transferred and rapidly reacts only at the regions where the raised part of the soft stamp makes contact with the surface (Jung et al., 2001; Kane et al., 1999). This technique allows patterns of hundreds of square centimetres to be created. The regions of unreacted gold that remain after the initial printing step can be reacted with a different SAM by immersing the entire assembly in an ethanol solution containing a different thiol (Jung et al., 2001). The same basic strategy can be used for other types of `inks' including alkylsiloxanes (Jeon et al., 1995). Microprinting can be used to pattern a variety of different surfaces including both gold and silver films. Gold is, however, preferable to silver for life science applications since silver is cytotoxic whilst gold is more bioinert. In order to prepare a suitable surface for microcontact printing, a glass coverslip or polished silicon wafer is evaporation-coated with a thin layer of titanium (1±5 nm thick). This layer acts as a primer for the subsequent gold layer. A thin layer of gold (10±200 nm thick) can then be evaporated onto the titanium (Takayama et al., 2000). SAMs of alkanethiols of gold are formed by the direct contact between the alkanethiol (`ink') and the gold surface. The surface properties and hence its ability to interact or not with biological molecules of the particular SAM are determined by its terminal groups (Ostuni et al., 1999). Thus the purpose of the SAM is either to facilitate the adsorption of cell adhesion proteins or totally to prevent adsorption depending upon the particular pattern's requirement. Patterned SAMs created using the microcontact printing technique are stable in cell culture medium at 37 ëC for several days. This stability is essential if, for example, detailed serial time studies are to be performed using the patterned substrates (Endler et al., 2003). Self-assembly monolayers (SAMs) Self-assembly molecules (SAMs) provide a means to produce bespoke surfaces. The original self-assembly monolayers (SAMS), disulphides on gold surfaces, were discovered in 1983 followed by alkanethioles shortly after. Today SAMs are used in a wide variety of applications, including the fabrication of sensors, transducers, protective layers including lubricants and as a means of patterning substrates. SAMs are the result of a number of molecules with an organic thiol group reacting with a thin layer of gold, silver or copper to create an interface film approximately 2 nm thick (Whitesides, 2003). Highly ordered SAM structures can be formed from long chain alkanethiols, HS(CH2)nX, where n is in the range 11±18 and X is a small, non-polar organic functional group

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(Whitesides et al., 2001). The sulphur atoms co-ordinate to the gold and the trans-extended alkyl chains pack tightly together, thus the terminal X group is confined to the interface between the body of the SAM and the aqueous phase. Therefore, the terminal X group largely controls the overall properties of the patterned surface (Mrksich et al., 1996). Using different X groups or chemically modifying the X group post SAM fabrication, it is possible to generate well defined surfaces with a broad range of characteristics. These range from being highly favourable for protein adsorption to the surface, to being almost totally resistant to protein adsorption (inert surfaces) (Whitesides et al., 2001). Thus, it is possible to indirectly (i.e. via adsorbed proteins) control the interaction of cells with a surface at the molecular level. The ability to form inert surfaces, i.e., surfaces which resist protein adsorption and therefore cannot support cell attachment, is particularly important in the generation of patterned surfaces. Terminating the SAM with a functional group composed of oligo(ethylene glycol) is one of the commonest methods employed to provide a near inert surface (Ostuni et al., 1999). This technique effectively makes it possible in two dimensions greatly to reduce protein adsorption, thus providing a cell-free zone against a well demarcated area of cell attachment. A good example of the use of two alkanethioles is provided by Mrksich et al. (1996). An outline of their experiment is now described. PDMS moulds having micrometre-scale relief patterns on their surfaces were fabricated using replica moulding to form a contoured film of polyurethane on a glass slide. Evaporation of a thin (<12 nm) film of gold onto the polyurethane was followed by the contact printing of hexadecanethiol [HS(CH2)15CH3]. Subsequently, a second SAM terminated in tri(ethylene glycol) groups was formed on the bare gold remaining by immersing the entire substrate in a solution of a second alkanethiol [HS(CH2)(OCH2CH2)3OH]. When this twocomponent pattern was immersed in fibronectin solution, the protein adsorbed only on the methyl-terminated plateau regions of the substrate and not on the oligo(ethylene glycol) interconnecting regions. Finally, bovine endothelial cells were added which attached only to the fibronectin covered regions. SAMs terminated in methyl groups are hydrophobic and adsorb proteins very quickly and irreversibly from solution. (Mrksich et al., 1996). Because the gold film is extremely thin, the resultant substrates are optically transparent which allows the attached cells to be visualised using, for example, a light microscope (Mrksich et al., 1997). Detailed reviews on SAMs of alkanethiolates have been written by Ostuni et al. (1999) and also by Takayama et al. (2000). Both accounts include elaborate examples of different SAMs and substrates. A typical example of a practical application of the process has been described by Endler et al. (2003) who developed a technique to characterise the in-vitro propagation of viruses using precisely patterned cultures of mammalian cells as hosts.

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Whilst SAMs have definite potential in many areas of cell patterning, they have a drawback in the field of tissue engineering, i.e., creating tissues in vitro. Some areas of tissue engineering, e.g., the bioengineering of complex human tissues, potentially require similar patterning precision to that provided by SAMs. The spatial control is required to direct cell growth for cell types that require precise spatial development to function (e.g. nerve cells), also to provide a mechanism for facilitating normal in-vivo cell-to-cell interactions. Alternatives to the use of SAMs have been devised. For example, Patel et al. (1998) have described a generic technology by which any biotinylated ligand can be patterned at the micron-scale onto the surface of a biodegradable polymer, e.g., polylactidepoly(ethylene glycol). Patel et al. claim that `the ability to pattern a wide range of ligands opens new possibilities in spatially controlled cell engineering, because the ability to control cell organization on the templates becomes limited by our understanding of the molecular biology of ligand-to-receptor binding and its influence on cellular development, not by surface engineering constraints'. Using their platform technology, spatial control of both bovine endothelial cells and nerve cells was demonstrated. Furthermore, directional control over neurite extensions from the nerve cells was achieved by the deployment of IKVAV peptide (isoleucine-lysine-valine-alanine-valine) in precise patterns. Another alternative to the use of SAMs is chemical micropatterning using low-temperature gas plasmas on fluorine containing polymers. Plasma treatment can create cell-friendly surface modifications without changing the bulk of the material, e.g., allowing the attachment of fibroblasts to PEEK (polyetheretherketone) (Ohl et al., 1999; Schroder et al., 2000). Microfluidic patterning using microchannels An alternative to microcontact printing using a PDMS stamp is to bring a stamplike form into direct contact with a surface to be patterned and retain it in place. The stamp-like form is fabricated by moulding it from a master which possesses a patterned network of surface reliefs that define the eventual areas of contact between the stamp-like form and the surface to be patterned. Thus, areas where the stamp-like form does not make direct contact with the surface creates a series of microchannels allowing liquid reactants to be guided accurately over the surface in a pattern defined by the stamp-like form (Delamarche et al., 1998). Using these channels, fluid can be accurately directed over the surface, since where the PDMS makes direct contact with the surface, fluid flow is absent. For this technology, the elastomeric properties of the PDMS has two big advantages. Firstly, provided there is conformal contact, a liquid-tight seal between the stamp and the surface can be created and maintained, thus creating the potential to form deep channels (>25 microns) (Folch et al., 1999). However, achieving and maintaining conformal contact is not a trivial matter. Secondly, the surface to be patterned need not be planar due to the flexibility of the PDMS

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form. Finally, it is worth noting that the microchannels have a very high surface:volume ratio which can be both beneficial or detrimental depending upon the particular application. By flowing cell adhesion proteins or ligands through the microchannels these components are deposited resulting in a patterned surface favourable for subsequent cell attachment. A good example of the technique is by Folch and Toner (1998) who created a network of deep channels by replica moulding PDMS using microfabricated glass master moulds. Microcapillaries were created by the self-sealing action of the PDMS microstructure against the surface of a chosen substrate. Submillilitre volumes of protein were injected into these microcapillaries. The protein adsorbed only onto areas of the substrate which were exposed to the microflow. The capillaries were flushed and the PDMS microstructure carefully peeled away. Both rat collagen and human plasma fibrinogen were deployed. For cell seeding, hepatocytes were chosen since they were known not to be able to attach to the bare substrate but would readily attach to either collagen or fibronectin. Thus selective attachment of hepatocytes was achieved. Interesting, after 24 hours the hepatocytes start spreading out from the patterned areas. This was attributed as possibly being due to the adsorption of endogenously secreted ECM protein by the cells (Odenthal et al., 1992). Finally, Folch and Toner (1998), created micropatterned cocultures of hepatocytes and fibroblasts by first adding sufficient hepatocytes in solution to totally cover the protein and then adding fibroblasts in solution which readily attached to the bare substrate. As an alternative to using a protein intermediate step and if the channels are large enough, it is possible to directly flow cells through the deep microchannels and thus pattern a cell-friendly surface in a more direct manner (Folch and Toner, 1998). Folch et al. (1999) have fabricated deep (>25 micron) channels by replica moulding PDMS using microfabricated glass master moulds. Fibroblasts in solution were directly injected into the microcapillaries created by the PDMS microstructure conformally sealed against the surface of a conventional tissue culture dish. The PDMS microstructure was carefully peeled off the cell-friendly surface after two hours leaving cells arranged in the desired pattern. A microchannel network can be filled simply by placing a drop of solution at the entrance of the primary channel (service port) and capillary attraction results in the fluid being drawn into the system. However, more sophisticated hydrodynamic methods can also be used if desired including syringe drivers and peristaltic pumps (Juncker et al., 2002). The range of surfaces suitable for micropatterning using microfluidic deposition is much greater than for microcontact printing and includes gold, glass and polystyrene (Delamarche et al., 1997). Typically, microfluidics networks require submicrolitre volumes of liquids resulting in excellent economy of the reagents (Folch and Toner, 1998). This is essential as the reagents for life science microfluidic experiments are frequently extremely precious (Delarmarche et al., 1998).

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A major challenge in microfluidics systems is to create three-dimensional arrangements. To date, the complexity of fabricating such structures has prevented their construction, although this situation may be changing through the deployment of rapid prototyping technologies. For example, Chiu et al. (2000) based on the work of Anderson et al. (2000) have fabricated three-dimensional microfluidic systems and used them to pattern proteins and mammalian cells onto a surface, e.g., bovine endothelial cells. Recently, Tan and Desai (2003a,b) have also reported favourable results using microfluidic patterning to create three-dimensional biopolymer matrices of collagen-chitosan-fibronectin seeded with endothelial cells and fibroblasts. Laminar flow patterning A variant of patterning using microfluidic networks is laminar flow patterning. The principle underlying this technique is that the flow of fluids passing through the microchannels is laminar due to the small bore diameter (typically <250 microns) and slow flow rates (maximum 1 cm per second). Such flows are described as having a low Reynolds Number (Re). This results in an interesting phenomenon when two or more microchannels are combined via a carefully designed junction into a single microchannel. If all the Re values are low, then the resulting stream will be composed of all the original components all flowing parallel to one another, i.e., there is no turbulent mixing. Mixing can occur only by the much slower process of diffusion (Fick's Law). Multilayered streams can be used for a number of tasks including patterning stripes of different reagents, e.g., cell adhesion proteins, cells and bioactive molecules. In 1999 Takayama et al. described a collection of results, including the patterning of surfaces with proteins, patterning of different types of cells (chicken erythrocytes and E. coli) adjacent to one another, the patterned delivery of reagents to pre-adhered cells and performing enzymatic reactions with selected cells (detachment of pre-adhered bovine adrenal capillary endothelial cells with trypsin/EDTA). Takayama et al. believe that laminar flow patterning has several key advantages over other cell patterning techniques, including (i) deploying easily generated multiphase laminar flows to pattern fluids and to deliver both reagents and cells for patterning, (ii) experimental simplicity, i.e., multicomponent patterns can be made without the need for multiple stages of pattern transfer which require precise registration, (iii) patterns can be created directly over living cells (impossible using other existing patterning techniques), (iv) patterning of the actual growth medium itself (again impossible using other current patterning techniques). The technique does, however, have limitations including the ability to generate only patterns of parallel stripes. Overall, multilayered stream technology has been suggested to have potential in a variety of applications including studies of cell-surface adhesion, chemotaxis, haptotaxis, cell-cell communication, and cellular ecology (Takayama et al.,

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1999). The ability to be able to co-culture cells is central to the fabrication of tissues and organs (Bhatia et al., 1999). Stencil patterning The simplest way to pattern a surface is to use a stencil. A stencil is a sheet of material with perforations which a reagent can pass through in order to create a printed pattern. Early stencils to produce cellular patterns were made of metal, including nickel stencils to deposit human amnion cells controllably and stainless steel stencils to pattern rat embryo neurons in order to create simple neural networks (Carter, 1967; Jimbo et al., 1993). However, since metal stencils cannot easily form a liquid-tight seal between the stencil and the underlying surface, liquids can easily seep into the adjacent masked-off areas where it can then deposit its proteins, ligands, etc. Furthermore, it is extremely difficult to fabricate a metal stencil at the micron-scale level. Because of these inherent problems, stencil design moved away from hard to soft stencils, e.g., PDMS. PDMS has the advantage of being capable of being accurately moulded at the micron-scale level plus it seals spontaneously on making conformal contact with a surface. Folch et al. (2000) have described the low cost fabrication of PDMS stencils and their use in accurately patterning primary rat hepatocytes, mouse fibroblasts and human neonatal foreskin epithelial cells. In their experiments, polystyrene surfaces were uniformly coated with collagen type I which was allowed to gel prior to applying the stencil. A solution of cells was then poured over the stencil. Once the cells had attached to the areas left exposed by the stencil, the stencil was gently peeled away from the surface. Whilst this technique can produce cell patterning, the level of resolution and the range of pattern flexibility is poor compared to microcontact printing. Folch et al. explored the limits of their patterning technique. In practice, resolution depended upon the height to width ratio of the perforations in the stencil, e.g., for human neonatal foreskin epithelial cells the ratio needed to be greater than 2.5 for patterning to be achieved. A 100 micron thick stencil perforated with rows of 40 micron diameter holes resulted in single cell micropatterns, however, when the hole diameter was reduced to 35 microns the pattern failed to materialise (Folch et al., 2000). Other PDMS variants for stencil production, e.g., using PDMS as a dry resist or dry lift off procedures have been described by Jackman et al. (1999) These procedures are termed, `dry' by the authors because the PDMS membrane conforms and seals reversibly to the surfaces without the requirement for a solvent. Further, their separation from the surface likewise does not require the use of solvent, just a gentle peeling action. Using the `dry' techniques, microstructures as small as five microns were fabricated with a variety of different materials including a biological material (bovine carbonic anhydrase).

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17.4 Three-dimensional controlled cell deposition techniques Human tissues have evolved as three-dimensional aggregations of cells and associated intercellular matrix, acting together to perform one or more specific functions within the body. The creation of three-dimensional tissue in vitro is termed tissue engineering, and is a distinct departure from growing twodimensional cell monolayers in conventional cell culture flasks. The complexity alone is several orders of magnitude greater. Typically, the basic tools and reagents required for tissue engineering are cells, a pre-formed threedimensional polymer scaffold mimicking the shape of a desired tissue or organ and a bioreactor. Under sterile conditions, cells are seeded onto the scaffold to produce a three-dimensional construct which is perfused with culture medium whilst its environment (including sterility) is controlled by the bioreactor in which it is placed. Growth and maturation of the construct takes weeks/months since the cells need to replicate and produce extracellular matrix in order to become a functional tissue (Niklason et al., 1999). To date a number of different tissues have been grown using the basic method described above including skin (Curran and Plosker, 2002) blood vessel (McKee et al., 2003) and bladder (Atala, 2003). A few commercial products have also entered into the clinic having gained full FDA approval including Apligraf skin (Organogenesis, Canton, MA, USA). It is estimated that nearly one hundred thousand patients have already been treated with tissue-engineered products (Mason, 2003a). Whilst tissue engineering is progressing, there are, however, potential drawbacks to the typical current approach of using pre-formed scaffolds including limiting both the speed of production and the complexity of the organ that can be fabricated. The limitations include, firstly, an inability to assemble tissues and organs rapidly, currently, this takes weeks/months which adds substantially to costs, and secondly, the deployment of a temporary polymer scaffold (Mason, 2003b). The rationale for deploying a polymer scaffold is that it acts as a temporary artificial extracellular matrix whilst the regular in-vivo extracellular matrix (e.g. collagen and elastin) is synthesised by the cells. The artificial matrix must be fully replaced by the regular extracellular matrix in order to have a normal piece of tissue. The disadvantages of using a scaffold include difficulties in achieving homogeneous cell seeding, difficulties in arranging mixed cell populations in the correct architecture and finally the biggest challenge of adequately vascularising the neotissue (Cassell et al., 2002; Nasseri et al., 2003). Vascularisation is essential if an actual tissue (cf. a monolayer of cells) is to be adequately perfused when implanted surgically. Potential solutions for creating tissues without the need of a pre-formed polymer scaffold have been demonstrated including forming arteries using a

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mandrel to roll layers of different cells types (smooth muscle cells, fibroblasts) around in order to produce a hollow tube (L'Heureux et al., 1998), `cellfriendly' fabrication of three-dimensional constructs from solutions of homogeneously mixed cells and polymer (Mason et al., 2004) (both these techniques are limited to tubular structures not requiring a microvasculature, e.g., small diameter blood vessels), three-dimensional photopatterning of cell-hydrogels (Liu and Bhatia, 2002), and organ printing (Mironov et al., 2003).

17.4.1 Organ printing Organ printing is at a very early stage of development but may have the potential to overcome the difficulties of conventional tissue engineering especially in the fabrication of solid tissues, e.g., liver and kidney. It differs from current advanced three-dimensional scaffold fabrication using, for example, rapid prototyping technologies in that the scaffold and the cells are fabricated together into a construct in one single action rather than creating a complex threedimension scaffold and then seeding it with cells (Gomes et al., 2002). Accurately seeding complex pre-formed scaffolds is technically extremely challenging (if not impossible) if as is universal, more than one cell type is required to make the particular tissue or organ. (Kim et al., 1998; Sherwood et al., 2002; Hutmacher, 2001). Organ printing exploits the viscoelastic nature of cell aggregates (small collections of cells). Cell aggregates when positioned sufficiently close to one another will `flow' together and fuse. This interesting phenomenon in multicellular animals has been known for nearly a century (Wilson, 1907). Wilson demonstrated that dissociated cells of silicious sponges would rapidly combine to form syncytial masses that had the power to differentiate into new sponges. The sponges were cut up with scissors and passed through a cloth mesh with the constituent cells passing through the cloth's pores. Upon settling in a container of sea water, the cells initially had amoeboid activity and attached themselves to the underlying surface. However, within a few minutes, the cells quickly fused to one another to form minute balls of cells which conglomerated together to form larger masses. Later, Ruud and Spemann (1922) demonstrated the phenomenon in a more complex animal. Using very early salamander embryos (gastrula) and placing a weighted fine glass needle across the gastrula. They then observed the passage of the glass needle as it passed transversely through the gastrula. Commenting on this paper, Steinberg and Poole (1982) noted this arrangement to `a weighted wire passing through a block of ice'. The passage through the gastrula occurred so slowly that the cells of the gastrula responded to the action of the needle in the manner of viscous liquid droplets. The cells slowly separating below the needle and then coming together (coalescing) after it had passed. The gastrula was observed to be unaffected by this event.

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Combining this phenomenon of fusion with knowledge gained from developmental biology is allowing researchers to begin to pioneer organ printing as a new platform technology for the fabrication of three-dimensional tissues in vitro. Living embryonic tissues have been demonstrated to behave in a manner similar to viscoelastic fluids with well described flow and fusion behaviour. In suspension or on non-adhesive surfaces, cell aggregates have been observed to coalesce in the fashion of liquid droplets coalescing to form one big drop (Steinberg and Poole, 1982; Forgacs et al., 1998; Mironov et al., 2003). Recent research by Thompson et al. (2003) has demonstrated similar observations in chick embryos. An embryonic chick heart `ventricle' was cut into loops perpendicular to its long axis. The loops contained only myocardium, endocardium lining and intervening matrix. These loops were threaded onto segments of vitelline vein which had been obtained from another chick, i.e., similar to threading beads onto a necklace. This arrangement was perfused with supplemented culture medium and incubated. Isolated loops beat steadily for several days, however, if the loops were placed in direct contact with one another, the loops fused together in less than 24 hours to form a muscular tube (myotube) and beat as one tissue. This tissue fusion process underpins organ printing. Mironov sub-divides the technology into three sequential steps; pre-processing, processing and postprocessing (Mironov et al., 2003). Preprocessing is the acquisition of an accurate three-dimensional computer aided design (CAD) of the desired organ or tissue. This can be derived from various imaging techniques including magnetic resonance imaging (MRI) and computer aided tomography (CT). An excellent review of recent developments in computer aided tissue engineering can be found in Sun and Lal (2002). Processing involves taking the CAD blueprint and using it to instruct a printer capable of three-dimensional printing. Postprocessing refers to the care and perfusion of the neo-organ. This overall process has two main advantages over conventional tissue engineering employing scaffolds. Firstly, cell printing is automated and fully computer controlled, thus potentially capable of high-speed fabrication of tissues. Secondly, the process is not limited to two-dimensional patterns since it does not depend upon surface modification (Wilson and Boland, 2003). The technology is broadly based on personal computers that are programmed to drive highly modified inkjet type printers, plotters or `cell dispensers'. The `inks' are cells and biodegradable gels (Markwald, 2003). The particular gels are liquid at 20 ëC but gel at above 32 ëC, i.e., they set at normal body temperature (Gutowska et al., 2001). Promising early experiments with organ printing have been described by Wilson and Boland (2003), Mironov et al. (2003) and Warren (2004). Warren describes his particular device as a bespoke bioassembly tool (BAT) with a three-axis printing capability which is able accurately to deliver cells, biomolecules and other biomaterials via a series of capillaries and nozzles. In contrast, Wilson's and Mironov's devices are fabricated from regular

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commercial inkjet printers, e.g., Hewlett-Packard (Palo Alto, CA, USA) and Canon (Tokyo, Japan). Since collections of cells (aggregates) cannot fit through standard inkjet printer heads, appropriately sized heads had to be fabricated, consisting of larger, but in the same configuration, multiple independent activated piezo electric pumps each of which is supplied living cells suspended in solution via a capillary tube. The pumps then feed the cell suspension to the inkjet needle for printing to finally take place. Modified proprietary printer software and electrical hardware are used to control the print head output. Using this arrangement allows the print head to move in 50 micron increments over an area of approximately 57 cm by 25 cm. To date, both proteins (albumin, steptavidin, biotin) and whole cells (bovine endothelial cells) have been three-dimensionally printed. Endothelial cells have been printed using a final needle size (30-gauge) which produced drops with an average diameter of 150 microns and volume 1.5  10ÿ5 ml which by using cell `ink' at a concentration of 1  105 cells/ml contained an average of one or two cells per drop. Proteins were printed onto glass coverslips and the cells were printed onto either reconstituted basement membrane (Matrigel, BD Biosciences, San Jose, CA, USA) or collagen. The cells were printed in a line, incubated at 37 ëC for three days and then imaged. Wilson and Boland demonstrated that at three days in culture, the cells were still arranged in a simple linear pattern and that the majority (75%) were still alive. The time of cell death appears to have been after the cells passed through the print needle, due to the small drop volume and thus the potential to totally evaporate prior to placing the freshly printed construct into cell culture media. However, this is probably a minor problem since one of the advantages of the technology is that living cells can be kept hydrated throughout the deposition process by controlling the water content of the gel support. Boland has further demonstrated the ability to place cell aggregates (endothelial cells) in a three-dimensional thermosensitive gel (poly[n-isopropylacrylamide-co-2-(N,N-dimethylamino)-ethyl acrylate] copolymer) close enough for fusion to occur (Boland et al., 2003). Thus the pieces of the jigsaw for organ printing are beginning to emerge, however, it is still at a very early stage in development. A good, brief overview of organ printing and descriptions of the various companies and research groups involved including Sciperio Inc (Stillwater, OK, USA) and Bhatia's group at the University of California (San Diego, CA, USA) can be found in a recent review by Constans (2003).

17.5 Future trends Future trends for two-dimensional chip-based controlled cell deposition are likely to parallel those already witnessed in the semiconductor industry, i.e., increasing performance and chip specification matched by an ever falling cost of

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production as demand increases. The disposable chip cannot be that many years away. Already major semiconductor manufacturers are taking more than a passing interest in the basic technology, e.g., IBM (IBM Research, Ruschlikon, Switzerland) (Juncker et al., 2002; Delamarche et al., 2002). In comparison to cell chips, the technology of organ printing is much further from actual application. Applications for both two-dimensional and three-dimensional cell deposition technologies range from the simple biosensor to clinical grade transplant organs and tissue. One possible application, `the animal on a chip', however, encompasses a number of these applications into one device and is therefore a good illustration of a medium-term future potential for controlled cell deposition techniques. `Animal on a chip' technology is an attempt to improve toxicology evaluation and efficacy of prospective pharmaceutical products by replacing traditional animal experiments with microscale animal cell culture analogue (CCA) devices. A CCA effectively aims to mimic the essential tissues of an entire animal, e.g., lung and liver on one chip with the individual tissues connected in series by microchannels through which flows substitute `blood', i.e., cell culture medium (Shuler et al., 1996). It is thought that a CCA system when used in conjunction with a physiologically based pharmacokinetic model, can possibly act as a human surrogate for predicting human responses to drugs (Sin et al., 2001). Shuler's group at Cornell University (Ithaca, NY, USA) has fabricated such a device consisting of three chambers each containing a different tissue relevant to drug research, e.g., liver for toxicology studies (Sin et al., 2004a). Culture medium is circulated around the body of the `animal on a chip' using a self-priming microfluidics diaphragm micropump (Sin et al., 2004b). The advantage of a chip-based CCA is that it has the potential for parallel highthroughput screening at a much reduced cost to conventional animal studies. Furthermore, the system requires only very small quantities of candidate liquids and since these candidate compounds are extremely expensive to synthesise thus further reduces the cost. Both organ printing and `animal on a chip' technology demonstrate the beginnings of a natural convergence between chip microfabrication technologies and tissue engineering, a convergence that is certain to continue at an everincreasing pace as the vital controlled cell deposition techniques continue to be invented and refined.

17.6 Further information Organisations, books and journals on both two-dimensional (chip-based) and three-dimensional (tissue-engineering-based) control cell deposition which may be of interest are listed below:

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17.6.1 Organisations Tissue Engineering Society International (TESi) Contact: Sarah Wilburn, TESi Administrator, 15 Arlen Road, Apt. J, Baltimore, MD 21236, USA. ([email protected]) Website: http://tesinternational.org European Tissue Engineering Society (ETES) Contact: Professor Robert Brown, ETES Secretary/Treasurer, University College London, Institute of Orthopaedics, Brockley Hill Stanmore HA7 4LP, UK. ([email protected]) Website: http://etes.tissue-engineering.net Pittsburgh Tissue Engineering Initiative (PTEI) Contact: David Smith, PTEI Secretary, 100 Technology Drive, 2nd floor, Pittsburgh, PA 15219, USA. ([email protected]) Website: http://www.ptei.org

17.6.2 Books (chapter of specific interest in parentheses) Lanza R P, Langer R, Vacanti V (2000), Principles of Tissue Engineering, San Diego California, USA, Academic Press, (Chapter 18 ± `Patterning of cells and their environment', Takayama S, Chapman RG, Kane R S Whitesides G M, 209±219). Palsson B, Hubbell J A, Plonsey R, Bronzino J (2003), Tissue Engineering, Boca Raton, Florida, USA, CRC Press, (Chapter 10 ± `Biomaterials: protein-surface interactions', Chinn J A, Slack S M, 10.1±10.13). Palsson B O, Bhatia S N (2004), Tissue Engineering, Upper Saddle River, New Jersey, USA, Pearson Prentice Hall, (Chapter 16 ± `Tailoring Biomaterials', 270±287).

17.6.3 Journals Biomedical microdevices Publisher: Kluwer Academic Publishers, Van Godewijckstraat 30, P.O. Box 17, 3300 AA Dordrecht, The Netherlands ± http://www.wkap.nl Lab on a chip Publisher: Royal Society of Chemistry, Burlington House, Piccadilly, London, W1J 0BA, UK - http://www.rsc.org

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Tissue engineering Publisher: Mary Ann Liebert Inc, 2 Madison Avenue, Larchmont, NY 10538, USA ± http://liebertpub.com

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Proceedings of SPIE ± Microfluidics and BioMEMS, 4650, 98±101. Sin A, Chin K C, Jamil M F, Kostov Y, Rao G, Shuler M L (2004a), `The design and fabrication of three-chamber microscale cell culture analog devices with integrated dissolved oxygen sensors', Biotechnology Progress, 20, 338±345. Sin A, Reardon C F, Shuler M L (2004b), `A self-priming microfluidic diaphragm pump capable of recirculation fabricated by combining soft lithography and traditional machining', Biotechnology and Bioengineering, 85(3), 359±363. Singhvi R, Kumar A, Lopez G P, Stephanopoulos G N, Wang D I C, Whitesides G M, Ingber D (1994), `Engineering cell shape and function', Science, 264 (5159), 696± 698. Steinberg M S, Poole T J (1982), `Liquid Behaviour of embryonic tissue', in Bellairs R, Curtess D, Dunn G, Cell behaviour. A tribute to Micheal Abercrombie, Cambridge, UK, Cambridge University Press. Stock U A, Vacanti J P (2001), `Tissue engineering: Current state and prospects', Annual Review of Medicine, 52, 443±451. Sun W, Lal P (2002), `Recent development on computer aided tissue engineering ± a review', Computer Methods & Programs in Biomedicine, 67 (2), 85±103. Takayama S, McDonald J C, Ostuni E, Liang M N, Kenis P J A, Ismagilov R F, Whitesides G M (1999), `Patterning cells and their environments using multiple laminar fluid flows in capillary networks', Proceedings of the National Academy of Sciences of the United States of America, 96 (10), 5545±5548. Takayama S, Chapman R G, Kane R S, Whiteside G M (2000), `Patterning of cells and their environment', in Lanza R P, Langer R, Vacanti J, Principles of Tissue Engineering, San Diego, USA, Academic Press, 209±220. Tan W, Desai TA (2003a), `Microfluidic patterning of cells in extracellular matrix biopolymers: Effects of channel size, cell type, and matrix composition on pattern integrity', Tissue Engineering, 9(2) 255±267. Tan W, Desai T A (2003b), `Layer-by-layer microfluidics for biomimetic threedimensional structures', Biomaterials, 25, 1355±1364. Thompson R P, Reckova M, de Almeida A, Bigelow M R, Stanley C P, Spruill J B, Trusk T T, Sedmera D (2003), `The oldest, toughest cells in the heart', Novartis Foundation Symposium: Development of the cardiac conduction system, Chichester, UK, Wiley, 157±176. Voger EA (1993), `Interface chemistry in biomaterial science', in Berg J C, Wettability (Surfactant Science Series 49), New York, USA, Marcel Dekker, 183±250. Vroman L (1974), `Surface charge, protein adsorption, and thrombosis', Science, 184 (136), 585±586. Vroman L (1987), `Methods of investigating protein interactions on artificial and natural surfaces', Annals of the New York Academy of Sciences, 516, 300±305. Vroman L, Adams A L (1986), `Rapid identification of proteins on flat surfaces, using antibody-coated metal oxide suspensions', Journal of Immunological Methods, 93 (2), 213±216. Warren W L (2004), `Digital printing of viable 3D-engineered tissue constructs', Presentation ± Engineering Tissue Growth Conference, Pittsburgh, USA, March 29± 31. Whitesides GM (2003), `The ``right'' size in nanobiotechnology', Nature Biotechnology, 21(10), 1161±1165. Whitesides G M, Ostuni E, Takayama S, Jiang X, Ingber D E (2001), `Soft lithography in

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biology and biochemistry', Annual Review of Biomedical Engineering, 3, 335±73. Wilson HV (1907), `On some phenomena of coalescence an degeneration in sponges', Journal of Experimental Zoology, 5, 245±258. Wilson Jr W C, Boland T (2003), `Cell and organ printing 1: Protein and cell printers. Anatomical Record Part A', Discoveries in Molecular Cellular & Evolutionary Biology, 272 (2), 491±496. Xia Y, Kim E, Zhao X M, Rogers J A, Prentiss M, Whitesides G M (1996), `Complex optical surfaces formed by replica molding against elastomeric masters', Science, 273 (5273), 347±349.

18

Biofouling in membrane separation systems Z C U I and Y W A N , University of Oxford, UK

18.1 Introduction A membrane can be defined as a barrier between two phases, which is permeable to some of the species present in these phases. It can be solid, liquid or gas, and membranes in the form of solid are most widely used. The heart of a successful membrane system is usually the membrane itself. Most membranes used are polymeric in nature, although inorganic membranes are becoming increasingly popular now. A wide variety of polymers with different separation properties have been used for membrane preparation, including cellulose esters (e.g., cellulose acetate (CA), regenerated cellulose), polysulfone (PS), polyethersulfone (PES), polyvinylidenefluoride (PVDF), polyethylene (PE), aromatic polymers (e.g., polyamide) or polyacrylonitrile, whilst some inorganic materials such as -Alumina/-Alumina, borosilicate glass, pyrolysed carbon, zirconia/stainless steel or zirconia/carbon have also been used to produce inorganic membranes. Membranes can be fabricated to form flat-sheet, tubular, hollow fibre or spiral wound modules. The driving force for membrane separation, related to biological and medical applications, is either pressure or concentration gradients. Dialysis is driven by the concentration difference (or the chemical potential difference) across the membrane. Dialysis is more useful for the removal/addition/exchange of small molecules between two aqueous phases, as the diffusivity of macromolecules is small and hence diffusion rate is too slow to have practical significance. Pressure driving processes are more practical. In terms of the separated species, pressure driven membrane separation processes can be broadly classified into four categories; reverse osmosis (RO), nanofiltration (NF), ultrafiltration (UF) and microfiltration (MF). Table 18.1 summarises the characteristics of these membrane processes. RO membranes are usually thought to be dense membranes without pores, which ensure more than 95±99% rejection of inorganic salts and charged organics. Solvent, usually water, goes through the membrane by solution and diffusion. NF is typically referred to as `loose' or `leaky' RO, and the NF

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Table 18.1 Characteristics of pressure-driven membrane processes Process

Mechanism

Membrane type

Size of rejected solute

Microfiltration Ultrafiltration Nanofiltration Reverse osmosis

Sieving Sieving, charge Sieving, charge Solution-diffusion

Porous Porous Porous, charged Dense/porous

80±10,000 nm 1±100 nm 0.5±5 nm <1 nm

membrane, usually regarded as porous with pore size in the range of 0.5±5 nm,* allows more salt passage through the membrane. UF membranes, mostly asymmetric, are capable of retaining species in the range of 300±500,000 Daltons of Ê (1±100 nm), while molecular weight, with pore sizes ranging from 10±1000 A microfiltration membranes are normally symmetric and have pore sizes of 80± 10,000 nm. These processes have been developed for separating particles or solutes with different dimensions. Reverse osmosis and nanofiltration can be used for the separation of salts and/or organics, ultrafiltration is particularly suitable for the separation of macromolecules such as proteins, while microfiltration is for the separation of micro-sized particles, e.g., cells and cell debris, from fluids. Today, microfiltration, ultrafiltration and nanofiltration are all commonly used membrane processes in biotechnological applications. Large-scale commercial uses of membrane separations have replaced conventional separation processes, and more are expected in the future. However, a problem hardly avoidable in all pressure-driven membrane system, known as membrane fouling ± a progressive decline in flux and a change in membrane selectivity, remains a main obstacle to widespread applications of membrane separations. In high-density cell cultures, the corresponding increase in the level of cell debris may cause unacceptable fouling of membrane systems (Belfort et al., 1994). Sampling systems using microfilters or ultrafilters guarantee bioprocess sterility and can be adapted on miniaturised bioreactors, showing great potentials for biological applications. However, biofilm growth on filter surfaces remains a problem and needs to be tackled properly (Gastrock et al., 2001). Fractionation of proteins using ultrafiltration is a very promising application of ultrafiltration in protein purification, which has attracted enormous interests in recent years (Higuchi et al., 1991; Saksena and Zydney, 1994; van Eijndhoven et al., 1995; Balakrishnan and Agarwal, 1996a; Li et al., 1997; Ghosh and Cui, 1998, 2000a,b; Wan et al., 2002), again fouling is a key factor affecting the productivity and selectivity of the process. While for membrane-based

* It should be noted that it is arguable whether pores exist in NF membranes particularly in the small `pore' size end.

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haemodialysis (artificial kidney) and artificial liver, in addition to flux decline and change in selectivity, fouling may also affect the biocompatibility of the membrane. For instance, plasma protein adsorption can cause not only the activation of different defence systems in blood, such as coagulation, complement and fibrinolysis, but also the adhesion and activation of blood cells (Vanholder, 1992; Basmadjian et al., 1997; Fujimori et al., 1998). The main focus of this mini review is to describe recent progress in understanding of membrane fouling in biological and medical applications, with emphasis on strategies of fouling control. A description of membrane separation is presented briefly with definitions, terminologies and applications. A discussion of less well understood fouling mechanisms and less accurate mathematical modelling for fouling, along with extensively examined affecting factors, is also included.

18.2 Membrane separation ± concepts and applications Normally, membrane processes achieve separation without change of phase and the operating conditions are very mild. When compared with conventional separation processes, membrane separation processes are often more capital and energy efficient, and in some cases, membrane processes can achieve totally novel results (Ho and Sirkar, 1992). In the following sections, concepts and terminologies related to membrane fouling and applications are briefly presented.

18.2.1 Nominal molecular weight cut-off (MWCO) The pore size of UF and NF membranes are characterised by their nominal molecular weight cut-off (MWCO), which is defined as the smallest molecular weight species for which the membrane has more than 90% rejection. MF membranes are characterised directly by the mean pore size (microns).

18.2.2 Permeate flux The productivity of the membrane can be described by permeate flux, which is the quantity of permeated liquid (kg or litre) per unit membrane area (m2) and unit time (h or s). The volumetric flux (Jv ) is defined as the permeate volume/ membrane area and unit time, and is given by Jv ˆ


p ÿ
 …Rm ‡ Rf †

18:1

where Rm is the resistance of the membrane, Rf is the resistance of the fouling layer,
P is hydrostatic pressure difference,
 is the osmotic pressure

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difference between feed and permeate phase as a result of solute rejection and  is the viscosity of solvent. In industrial applications, Jv is often described by litre per square metre per hour (L/m2 h), although in SI unit, Jv has the unit m/s. It is clear that permeate flux is proportional to the applied pressure, reduced by the osmotic pressure difference
. The higher the concentration difference, the higher the osmotic pressure difference (
). In the MF and UF of macromolecules, the osmotic pressure is small and can be neglected. However, this should be checked for each application since it can result in significant error even in the separation of macromolecules.

18.2.3 Rejection The observed or apparent solute rejection coefficient is defined as R ˆ …Cb ÿ Cp †=Cb ˆ 1 ÿ Cp =Cb

18:2

where Cp is the concentration of the solute in the permeate, and Cb the concentration in the bulk or retentate. When Cp ˆ 0, only solvent, i.e., water passes through the membrane and R ˆ 1, and the solute is completely retained. If Cp ˆ Cb , the solute in the feed passes through the membrane freely and R ˆ 0. There is no rejection at all. Normally, 0 < R < 1, the value of R indicates the selectivity of the membrane for certain solutes. Another parameter used to express the solute rejection is the intrinsic rejection coefficient Ri , which is defined as Ri ˆ 1 ÿ Cp =Cw

18:3

where Cw is the concentration at the membrane surface. Ri is a property of the membrane and independent of hydrodynamic conditions. However, Ri is not normally used because the concentration of solute at the membrane surface is difficult to determine.

18.2.4 Transmission Solute transmission is another commonly used parameter in ultrafiltration. The observed transmission, S, is defined as the ratio of the solute concentration in the permeate, Cp , to that in the bulk or retentate, Cb . S ˆ Cp =Cb ˆ 1 ÿ R

18:4

As with the intrinsic rejection coefficient Ri, the intrinsic transmission coefficient Si is defined as the ratio of the solute concentration in the permeate, Cp, to that in immediately upstream of the membrane, Cw. Si ˆ Cp =Cw

18:5

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18.2.5 Membrane selectivity For the separation of a solution containing two solutes, it may be desirable to achieve maximum transmission for one component and minimum transmission for the other. The efficiency of a fractionation process can be expressed in terms of membrane selectivity, which is defined as: ' ˆ S1 =S2 ˆ …1 ÿ R1 †=…1 ÿ R2 †

18:6

where S1 is the transmission of one solute and S2 is the transmission of the other solute, while R1 and R2 refer to the retention of the two solutes respectively.

18.2.6 Concentration polarisation As water and low molecular weight components pass through the membrane, while high molecular weight solutes are retained, a concentration gradient builds up between the membrane surface and the bulk fluid. This accumulation of solute molecules or particles on the membrane surface is called concentration polarisation (Fig. 18.1). Concentration polarisation is common to all pressure driven membrane processes. It leads to rapid flux decline at the early stage of filtration, as shown in Fig. 18.2. Concentration polarisation depends on permeate flux, and hence operating parameters, such as pressure, temperature, feed concentration or velocity, and is reversible. Therefore, the severity of concentration polarisation can be lessened by increasing the rate of solute mass transfer back away from the membrane (Zydney, 1996), e.g., using hydrodynamic management. Concentration polarisation can be analysed within the mass transfer boundary layer illustrated in Fig. 18.1 where  is the distance over which the concentration changes from Cb (bulk concentration) to Cw (concentration at membrane surface), D is the diffusion coefficient for solute transport in solvent. The

18.1 Concentration polarisation in cross-flow ultrafiltration.

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18.2 Flux decline in membrane filtration.

convective transport of solute by the permeate flow towards the membrane is balanced by the back diffusion to the bulk, i.e.,   dC Jv …C ÿ Cp † ˆ ÿD 18:7 dy Integrating eqn 18.7 across the boundary layer from 0 to , the flux is given by: Jv ˆ k ln ‰…Cw ÿ Cp †=…Cb ÿ Cp †Š

18:8

where k …ˆ D=† is the mass transfer coefficient. For total injection, i.e., Cp ˆ 0, then eqn 18.8 gives, Jv ˆ k ln …Cw =Cb †

18:9

18.2.7 Gel layer As the transmembrane pressure (TMP ˆ Pb ÿ Pp ) increases, the flux initially increases, and so does the wall concentration, Cw . When Cw reaches the gelation concentration of the macromolecules (Cg ), or the solubility of the rejected salt (as in NF), a gel or precipitate layer may form. In this case, Cw reaches a maximum and, the flux no longer increases with TMP. The operation is said to be in the pressure independent region (see Fig. 18.3). The pressure independent flux is called the limiting flux, which depends on feed flow rate, or mass transfer coefficient, as Jv ˆ k ln …Cg =Cb †

18:10

The phenomenon of a limiting flux can also occur in the absence of a physical gel. The increase in concentration at the membrane surface increases the osmotic pressure and decreases the mass transfer coefficient. These two effects alone can give a plateau in the flux-applied pressure curve (Field and Aimar, 1993). The

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18.3 Permeate flux vs. applied pressure plot for cross-flow ultrafiltration.

mass transfer coefficient in eqns 18.8, 18.9 and 18.10 can be estimated from correlations of the Sherwood number in terms of the Reynolds number and the Schmidt number. Detailed correlations have been summarised by Hwang and Kammermeyer (1975).

18.2.8 Fouling Fouling refers to the deposition of some solute components on the membrane surface, at its pore openings or within the membrane pores that are caused by specific physical and/or chemical interactions between the membrane and the solute. The rate and extent of membrane fouling in any given system depend on the strength of the intermolecular interactions between the solutes and the membrane, in combination with the effects of the various hydrodynamic (and body) forces acting on the solutes (Zydney, 1996). Membrane fouling is time dependent and partially dependent on operating parameters such as pressure, temperature, velocity and feed concentration. It typically manifests itself as a decline in permeate flux (see Fig. 18.2) and an alteration in membrane selectivity. Therefore, these changes often continue throughout the process, and eventually require extensive cleaning or replacement of the membrane. It must be noted that the long-term decline in permeate flux may also be caused by the membrane compaction over time. Fouling is an undesirable effect in membrane processes. However, in certain special cases, controlled fouling can be utilised to favourably alter some of the membrane surface properties such as membrane surface charges to improve membrane rejection characteristics (Ghosh and Cui, 1998).

18.2.9 Applications Ultrafiltration and microfiltration are currently used in a very wide range of applications in food, pharmaceutical, biomedical, biological, chemical, water

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and wastewater industries. The major areas of biological applications can be conveniently summarised as follows: 1.

2. 3.

4. 5. 6. 7.

8.

Concentration · protein (enzymes, milk, egg white, etc.) · starch and pectin · yeast production, mammalian cell harvest. Diafiltration · desalting and removal of other low molecular weight compounds · buffer exchange. Clarification · fruit juice and wine (removal of haze components) · beer (removal of cellular debris and bacteria) · sugar refining (removal of polysaccharide, proteins, colloidal impurities). Purification · fractionation of proteins · separation of antibiotics or vaccines from fermentation broths. Sterile filtration and virus removal for bioproducts. Biomedical uses · artificial organs (artificial kidney and liver) · tissue culture. Membrane bioreactors · antibody production · vaccine production · cell culture. Effluent treatment · domestic wastewater · industrial wastewater.

18.3 Fouling mechanisms and factors affecting fouling 18.3.1 Fouling mechanisms and mathematical description of fouling Fouling is complicated and characterised by a progressive decline in flux with time. Though having been extensively studied, membrane fouling remains a poorly understood phenomenon. The fouling process undergoes different stages at different timescales (Howell and Velicangil, 1980; Aimar et al., 1988). For example, Hallstrom et al. (1989) proposed a three-step process to describe fouling by protein. Firstly there is a rapid deposition in the first minute on the membrane surface and at the entrances to the pores, resulting in a sharp flux loss; further deposition (in the first hour) then occurs on the top of the first deposited

Biofouling in membrane separation systems

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layer, causing a slower increase in the membrane resistance (hence a slower flux decline) than the initial deposition; and thirdly, eventual bridging of the pore entrances occurs and a complete surface layer builds up, and this could be a quasi-steady-state period where flux decreases very slowly. Though the very nature of membrane fouling is not completely clear, considerable understanding can still be obtained by defining two classes of fouling (Davis, 1992): internal membrane fouling, which is caused by the deposition within the internal pores of the membrane or directly onto the membrane surface due to adsorption, precipitation, pore plugging, etc., and external cake fouling, which refers to the formation of a stagnant cake (or gel) layer on the membrane surface as a result of concentration polarisation. The external cake and the fouled membrane contribute two resistances to filtration in series. Both types of fouling generally increase with time, resulting in a flux decline with time. It should be pointed out that once the cake is formed, the filtration performance will be changed from membrane-dominated separation to fouling layer-dominated separation. Cake formation during membrane filtration has been demonstrated by scanning electron microscopy (SEM). In bovine serum albumin (BSA) filtration with a 0.16 m polyethersulphone membrane, the membrane surface was found to be completely covered by a BSA deposit after only 30 min of filtration (Kelly and Zydney, 1994). Similar results were reported by Tracey and Davis (1994) for BSA filtration through track-etched polycarbonate microfiltration membranes and by Kim et al. (1992) for BSA filtration through a variety of totally retentive and partially retentive ultrafiltration membranes. Atomic force microscopy (AFM) images also indicate that the membranes are totally covered by a protein layer after filtration of a 0.5 g/l BSA solution for 30 min (Huisman et al., 2000). More recently, James et al. (2003) used SEM, AFM and X-ray photoelectron spectroscopy (XPS) to comprehensively characterise the surface of microfiltration and ultrafiltration membranes. The foulant layer on the surface of membranes after milk filtration confirmed the formation of a gel layer (cake) on the membrane surface following 30 min filtration of skimmed milk. Generally, these deposits are about several microns thick (Kelly and Zydney, 1994; James et al., 2003). Protein deposits have also been identified within the membrane structure. Hanemaaijer et al. (1989) compared the amount of protein adsorbed on a nonporous polysulfone surface with that adsorbed on a membrane and found the former was 100±400 times less than the latter, suggesting that proteins adsorb within the membrane structure. GuÈell and Davis (1996) showed that fouling by BSA or lysozyme alone was dominated by pore blocking (internal fouling) during filtration using polysulfone and polycarbonate microfiltration membranes. Using immunochemical staining technique and transmission electron microscopy (TEM), Sheldon et al. (1991) demonstrated that BSA not only formed a cake layer at the membrane surface, but also accumulated within the membrane matrix of polysulfone and regenerated cellulose membranes. Su et al. (1998, 1999)

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employed small angle neutron scattering (SANS) to identify the location of fouling caused by BSA on ceramic membranes in situ under dynamic filtration conditions and found that the pore entrance of membrane is rapidly blocked at pH 3 (cake formation); while at pH 7, cake or gel layer formation is insignificant and fouling is mainly caused by the deposition of protein inside membrane pores. In the filtration of skimmed milk, analysis with XPS clearly indicated that protein was present within the membrane pore and the support layer, as well as on the skin layer of membrane (James et al., 2003). SEM images of a fouled membrane also clearly show that protein particles interact with the pore walls of the membrane (protein-polymer interaction) and these proteins also form agglomerates (protein-protein interaction) (James et al., 2003). As a result, the pores of a membrane can be narrowed and ultimately blocked. To describe membrane fouling in ultrafiltration, Hermia (1982) suggested that all fouling processes of ultrafiltration can be accounted for by using four theoretical kinetic models commonly employed for systems showing flux decline: complete blocking, intermediate blocking, standard filtration and cake filtration models. So-called `complete blocking' assumes that each fouling particle arriving at the membrane blocks some pore or pores without a superposition of particles. It leads to a time law for the permeate flux which is given by: ln Jv ˆ ln Jv;0 ÿ kb t

18:11

where t represents time, Jv;0 is the permeate flux per unit of area through the membrane at t ˆ 0, kb is the kinetic constant for complete blocking model. The other three models can be expressed through the following equation: Jv;0 ˆ …1 ‡ kf t†n Jv

18:12

where kf is the general kinetic constant for the fouling models. When n ˆ 1, it represents the so-called `intermediate blocking' model, which presumes that each particle can either settle on other particle previously arrived and already blocking some pores or it can directly block some membrane area. When n ˆ 2, it represents the `standard blocking' model, which means that each particle arriving at the membrane will deposit on the internal pore walls, leading to a decrease in the pore volume. When n ˆ 1=2, it represents the `cake filtration', which means that each particle will settle on another one previously arrived which has already blocked some pores and there is no room for a direct obstruction of any membrane area. For fouling in microfiltration, the cake filtration theory and `classical block laws' can be used (e.g. Hermans and BredeÂe, 1936; Grace, 1956):  n d2 t dt ˆ k 18:13 f dV 2 dV

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or dJv ˆ ÿkf Jv …Jv A0 †2ÿn 18:14 dt where t is the filtration time, V is the total filtrated volume, and Jv ˆ …1=A0 †dV =dt is the filtrate flux. The exponent n characterises the filtration model, with n ˆ 0 for cake filtration, n ˆ 1 for intermediate blocking, n ˆ 3=2 for standard blocking, and n ˆ 2 for complete pore blocking. As the blocking laws regard the membrane as a collection of parallel capillary tubes of constant diameter and length (Hermans and BredeÂe, 1936; Hermia, 1982), it would be invalid to apply these models to the fouling of polymeric MF membranes with highly interconnected pore structures (Ho and Zydney, 1999b). Membranes with an interconnected pore structure became fouled much more slowly than those with straight-through (nonconnected) pores (Ho and Zydney, 1999a,b). Therefore, there is a need for improved model development (e.g., Ho and Zydney, 2001). Mathematical modelling in conjunction with experimental research is a powerful tool in the study and understanding of membrane fouling. However, biological solutions are very complex systems, the predominant fouling mechanism is a function of the experimental conditions, depending on operating parameters, membranes characteristics and the properties of the feed solution. Moreover, protein molecules may change orientation and conformation during or after interaction with either the membrane surface or other protein molecules (Sheldon et al., 1991; Kelly and Zydney, 1994; Marshall et al., 1993). The development of more accurate and applicable models relies on much more detailed characterisation of the colloidal solutions to be filtered and membrane microstructure and surface properties.

18.3.2 Factors affecting fouling The rate and extent of membrane fouling depend on many factors, including membrane material (morphology, hydrophobicity), feed concentration and characteristics, solution environment (pH, ionic strength), and other operating parameters such as transmembrane pressure, shear rate, temperature (Fane and Fell, 1987; Matthiasson, 1983; Ko et al., 1993; Crozes et al., 1997; Babu and Gaikar, 2001). Electrokinetic effects (membrane and solute charge, pH and ionic strength), in particular, have been shown to have a significant effect on both fouling and retention (Ricq et al., 1999; Balakrishnan and Agarwal, 1996a). Membrane materials and membrane morphologies Matthiasson (1983) examined the relationship between the concentration of BSA solution, the amount of protein adsorbed onto different membrane surfaces,

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and the corresponding flux decline. For cellulose membranes, only monolayer adsorption was indicated, while for polysulfone or polyamide membranes, the protein adsorption increased and the permeability decreased as protein concentration was increased. Babu and Gaikar (2001) studied the BSA adsorption on cellulose triacetate (CTA) and regenerated cellulose (RC) membranes during ultrafiltration and found that the CTA membrane adsorbed more protein than the RC membrane, the fouling of the former was also much severer than the latter. Hydrophobicity of membrane materials is another important factor for membrane fouling. Protein normally adsorbs less to hydrophilic membranes than to hydrophobic membranes (Matthiasson, 1983; Hanemaaijer et al., 1989), and a hydrophilic membrane generally shows less fouling (Gekas and HallstroÈm, 1990; Babu and Gaikar, 2001). In the sterile filtration of recombinant human growth hormone, the addition of polysorbate 20 (0.05%) in the protein solution significantly reduced the fouling tendency due to the hydrophilic modification of membrane surface by polysorbate 20 (Maa and Hsu, 1998). Membrane morphologies can also affect the fouling performance during filtration. Ho and Zydney (1999a) examined the effect of membrane morphology and pore structure on protein fouling using different track-etched, isotropic, and asymmetric microfiltration membranes, and found that membranes with interconnected pores fouled more slowly than membranes with straight-through pores since the fluid could flow around the blocked pores through the interconnected pore structure. Similar results were also obtained in the filtration of BSA during normal flow (dead-end) microfiltration (Zydney and Ho, 2003). In filtration of a protein mixture (Ovalbumin + Lysozyme + BSA) with four different membranes having the same pore size (0.2 m) but different porosity, it was found that fouling behaviour was more affected by membrane morphology than by membrane surface chemistry (GuÈell and Davis, 1996). This is because the primary cause of flux decline is the deposition of protein aggregates (affected by morphology) rather than the absorption of single protein molecules (affected by surface chemistry) when the membrane pore sizes are much larger than the size of individual protein molecules (Jonsson and Johansen, 1991; MuÈller and Davis, 1996). In cell recovery and washing, cell recycling and cell debris removal by membrane filtration, it is generally found that UF membranes with a high MWCO (>100 kDa) give better flux than MF membranes in the long term (Defrise and Gekas, 1988), suggesting that membranes with a bigger pore size may be more affected by internal fouling. Solution environment The main parameters to characterise a feed solution include concentration, composition of the feed solution, pH, ionic strength and possible pretreatment. It is well accepted that these parameters can affect membrane fouling significantly.

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Increasing feed concentration generally results in a decrease in permeate flux. It has been observed in membrane filtration of both protein solutions (Matthiasson, 1983; Cheryan, 1986; Kuberkar and Davis, 1999; Su et al., 1999) and cell suspensions such as yeast (Patel et al., 1987), bacterium (Tejayadi and Cheryan, 1988) and mycelial culture (Haarstrick et al., 1991). This is a combined result of concentration polarisation (reversible, i.e., which will disappear if filtration stops) and membrane fouling. Membrane fouling, if initiated with adsorption like that in UF and MF, is directly linked to concentration polarisation, as the latter affects the quantity adsorbed at the membrane surface. The filtration of BSA using 0.07 m cellulose acetate membranes showed that higher BSA concentrations led to faster fouling. This fouling resulted in both lower overall transmission and lower permeate flux for higher BSA concentrations (Kuberkar and Davis, 1999). In a pilot-scale harvest process of recombinant yeast using ultrafiltration, Russotti et al. (1995) showed that higher initial cell concentration led to much faster flux decline and a higher fouling rate due to more mass deposition, and the apparent cake resistances also followed the trend with increasing initial cell concentration. Permeate flux can be improved by feed pretreatment, e.g., prior filtration or clarification of the feed solution (Lee and Merson, 1976; Belfort et al., 1994; Kelly et al., 1993). Kelly et al. (1993) demonstrated that the content of BSA aggregates could significantly affect the fouling performance of a 0.16 m PES membrane during BSA microfiltration. If BSA aggregates were removed by prefiltering the BSA solution through a 100 kDa UF membrane, the prefiltered BSA solution exhibited almost no flux decline over a 30 min filtration period. In the sterile filtration of human growth hormone, when the solution was pretreated with polysorbate 20, trace amounts of aggregates previously detected were not detectable any more, and the fouling tendency of the membrane could be reduced significantly (Maa and Hsu, 1998). It was suggested that the deposition of aggregated or denatured proteins may act as initiation sites for the continuous deposition of bulk proteins (Kelly et al., 1993). Wickramasinghe et al. (2002) found that flocculation of yeast suspensions prior to microfiltration using cationic polymeric flocculants could reduce fouling and enhance permeate flux. Sur and Cui (2004) demonstrated that the removal of extracellular proteins in a yeast suspension by diafiltration resulted in significant flux enhancement and reduced fouling rate during washed yeast microfiltration. Flocculation and washing both remove or aggregate smaller particles or macromolecules to form larger sized particles. Removal of extracellular proteins reduces the adsorption onto the surface of the membrane and the internal surface of the pores, and hence reduces fouling. Removal of smaller particles reduces the chance of pore plugging and internal fouling. pH and ionic strength have been found to have a profound effect on membrane fouling in protein filtration. The severest fouling in protein filtration

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usually occurs at the isoelectric point of the protein being filtered, and results in pronounced flux decline. At the same time, many authors have found that the protein transmission is highest when the solution pH is equal to its isoelectric point due to the fact that electrostatic repulsion is absent (Nystrom et al., 1998; Ricq et al., 1999; van Reis et al., 1997; Wan et al., 2002; Ghosh et al., 2003). Fane et al. (1983) studied the ultrafiltration of protein (BSA) solutions with retentive membranes over a range of pH values (2±10) and salt concentrations. A flux minimum occurred at pH 5 in the absence of salts, but in the presence of 0.2 M NaCl, flux increased monotonically with pH. Maximum protein adsorption occurred at the isoelectric points as well, and it was greater in the presence of salts. Conformational changes and charge properties of the BSA appeared to be the dominant factors determining the permeate flux. Oppenheim et al. (1996) examined the fouling of a 100 kDa MWCO polysulfone membrane by BSA. The BSA solutions were at pH 5 and 7 with 0.05 M and 0.15 M NaCl concentration. It was found that solution properties did not affect the amount of adsorbed material, but they did affect the resistance per mass of adsorbed species (i.e., the packing of molecules), with the maximum being at the isoelectric point of the protein. Huisman et al. (2000) studied the effect of pH on permeate flux, streaming potential and protein transmission during filtration of BSA solutions using various membranes with different MWCO. They also found that the lowest flux occurred at the isoelectric point of BSA. The protein transmission during a fouling experiment changed slowly from a value related to the ratio of protein size/pore size to the structure of the fouling layer. The significant effects of the solution pH and ionic strength on the cell adsorption onto a membrane surface (static) and membrane fouling during filtration have also been observed. In the ultrafiltration of Glutamicus, highest adsorption and severest fouling occurred at the isoelectric point of the cell, and the permeate flux decreased further upon the addition of an electrolyte into the feed solution at any pH away from its isoelectric point (Li and Fu, 2002). Meireles et al. (2003) examined the hydraulic resistance of cakes formed during the ultrafiltration of Streptomyces pristinaespiralis broths for different harvesting conditions. They found that hydraulic resistance associated with cake build-up was pH dependent. For broths at pH 2 or 3, hydraulic resistance associated with cake build-up was directly determined by interactions between the cells. While for broths at pH 4, hydraulic resistance associated with cake build-up was determined both by cell interactions and cell morphology. With ageing, cell surface interactions decreased due to the binding of a soluble component released by microorganisms, and as a result, specific hydraulic resistance increased accordingly. Sur and Cui (2005) observed the effect of pH on permeate flux in yeast microfiltration. When pH was increased from 3 to 5.5, the flux of microfiltration of 1% yeast dropped by 56%. This is obviously related to the electrostatic interactions within the cake layer.

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Operating conditions Operating conditions, including transmembrane pressure or permeate flux, shear rate or cross-flow velocity, can also affect membrane fouling during protein UF (Fane and Fell, 1987; Bowen et al., 1995; Nystrom et al., 1998; Ghosh, 2002; Velasco et al., 2003) and cell MF (Patel et al., 1987; Russotti et al., 1995). Generally, with the increase in transmembrane pressure, permeate flux increases as long as the flux is below the limiting flux. Aimar et al. (1988) showed that during the ultrafiltration of cheese whey with inorganic membranes, the initial flux decline with cheese whey became larger with an increase in transmembrane pressure, suggesting that severer fouling occurred at higher pressure. A similar trend was also found in the filtration of BSA using CTA and RC ultrafiltration membranes (Babu and Gaikar, 2001) and using negatively charged microfiltration membranes at pH 4, 5 and 6 (Velasco et al., 2003). Velasco et al. (2003) also showed that the fouling mechanism could change from internal fouling to external fouling as the applied pressure was increased, leading to higher final fluxes. At pH 4, this effect was more profound due to the formation of a loose protein network on the membrane surface. Marshall et al. (1993) reviewed the microfiltration of proteins and cells, and concluded that in microfiltration, increasing applied pressure could result in a much higher fouling rate, and in some cases the permeate could decline to less than the flux at lower pressure. Therefore, there is an optimum pressure at which the flux can be maximal. Similar results have been obtained by Krsti et al. (2001) for the microfiltration of Polyporus squamosus fermentation broth with a 0.2 m aluminum oxide membrane and by Sur and Cui (2001) for yeast MF. Castilho and Anspach (2003) studied the separation of mammalian cells, Chinese hamster ovary (CHO) and baby hamster kidney (BHK) cell lines using 0.45 m polysulfone membranes in a plate-and-cone dynamic filter. They found that flux increased linearly with increase in transmembrane pressure, and concluded that no significant membrane fouling or blocking occurred over the experimental period, given the low transmembrane pressure applied. This might be because the operations were carried out at fluxes below the critical flux, as fouling does not occur when the critical flux is not exceeded (Field et al., 1995). Increasing the shear rate, e.g., the stirring speed for stirred cell and the crossflow velocity in crossflow filtration and introducing flow instability, generally results in an enhancement of permeate flux and a reduction of fouling rate in both ultrafiltration (Aimar et al., 1988; Meireles et al., 1991; Kim et al., 1993) and microfiltration (Patel et al., 1987; Taddei et al., 1990; Tanaka et al., 1993, Bellhouse et al., 2001, Sur and Cui, 2001). Bowen et al. (1995) studied microfiltration of BSA solutions (1 and 0.1 g/l) and found that an increase in shear stress resulted in a decrease in membrane fouling. Patel et al. (1987) demonstrated that increasing crossflow velocity resulted in less rapid flux decline and a higher steady state flux for the microfiltration of a yeast medium.

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A similar trend was reported by Tanaka et al. (1993) for the filtration of a yeast suspension in saline. In pilot-scale crossflow microfiltration of a recombinant strain of Saccharomyces cerevisiae containing an intracellular product, Russotti et al. (1995) showed that increasing shear rate led to a higher overall average flux and lower specific cake resistance. It should be pointed out that most proteins and cells, particularly mammalian cells, are sensitive to shear. High shear or excessive exposure to shear may lead to protein denaturation, aggregation (Bowen and Gan, 1991, Kim et al., 1993; Persson and Gekas, 1994) and cell rupture or lysis (Vogel and Kroner, 1999; Castilho and Anspach, 2003), which in turn will cause severe fouling. Other operating parameters which can affect fouling include temperature and operation history. Normally, increasing temperature will increase permeate flux due to decreased viscosity and increased diffusivity (Meireles et al., 1991; Babu and Gaikar, 2001), again elevated temperature-induced protein denaturation or aggregation (Kelly et al., 1993; Persson and Gekas, 1994) should be avoided. In addition, there are reports showing that the history of the membrane used and the operation history can also affect its fouling and filtration performance (Ghosh, 2002; Li and Fu, 2002).

18.4 Biofouling Biofouling during filtration of biological solutions or suspensions is extremely complex due to the enormous range of biofoulants which may be present in any given aqueous system. This includes a variety of proteins, organic acids (e.g., humic acids), extracellular polymeric substances (EPS) secreted by bacteria, etc. In the following sections, fouling caused by proteins and microorganisms is discussed.

18.4.1 Fouling by proteins Protein fouling during microfiltration and ultrafiltration remains a very controversial topic, with considerable disagreement over both the mechanisms and rate of fouling as well as its overall importance relative to the concentration polarisation effects. Extensive reviews of previous work in this area have been presented by Nilsson (1990), Marshall et al. (1993) and Sablani et al. (2001). BSA is the most intensively studied model protein in both ultrafiltration and microfiltration. Bowen and Gan (1991) concluded that deposition of protein inside membrane pores (internal fouling) was an important mechanism for flux loss, whereas Kim et al. (1992) concluded that protein fouling of various ultrafiltration and microfiltration membranes was a surface phenomenon (external fouling), including fouling by multilayer (cake) coating and fouling by aggregates of proteins. They also suggested that aggregation might be

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initiated by rapid supersaturation of proteins above the pores due to high concentration polarisation. Opong and Zydney (1991) examined the effect of both BSA adsorption and deposition on the hydraulic resistance of several ultrafiltration and microfiltration membranes. They found that hydraulic resistances of clean 30 kDa (fully retentive membrane), 100 kDa and 0.16 m membranes were, as expected, significantly different, but the total hydraulic resistances of the three membranes after filtration of a 5 g/l BSA solution for four hours were actually quite similar (nearly two orders of magnitude increase in the resistance of the microfiltration membrane). However, some researchers showed that fouling was dependent on membrane materials and morphologies (Nystrom et al., 1998; MuÈller and Davis, 1996; GuÈell and Davis, 1996). MuÈller and Davis (1996) studied the fouling behaviour of different membranes in the filtration of BSA solution. Their results showed that 0.2 m track-etched polycarbonate (PC) membranes were internally fouled, with external fouling becoming predominant only at later stages; 0.2 m cellulose acetate (CA) membranes showed only internal fouling, while 0.2 m polysulfone (PS) and polyvinylidene fluoride (PVDF) membranes showed only external fouling. They also noticed that with increasing the feed concentration, resistance increased quickly and sharply, and external fouling became the dominant mechanism within 30 min for PC membranes. Ho and Zydney (2000) studied the transition from internal fouling to external fouling, and proposed a combined pore blockage and cake filtration model to describe the phenomenon. Although much of the research on protein fouling has focussed on the behaviour of BSA, a large number of practical proteins have also been examined. James et al. (2003) studied the ultrafiltration and microfiltration of skimmed milk and found that proteins could form a layer over the membrane surface and deposited within the pores as well. Bowen and Gan (1992) examined the fouling caused by enzyme yeast alcohol dehydrogenase (YADH) during microfiltration. It was observed that permeate flux decreased continuously with time. This decrease could be described using the standard blocking filtration law, indicating that fouling was caused by the deposition of enzyme on the walls of the pores (internal fouling) under the given experimental conditions. Chan et al. (2002) studied the ultrafiltration of binary protein mixtures containing globulin (GG), lysozyme (LYS), BSA, and -lactoglobulin (BLG). The respective protein species deposited onto membrane surfaces was carried out using matrix-assisted laser desorption ionisation mass spectroscopy (MALDIMS). The results indicated that much of the deposition was the transmitted species, probably from the surface, the pores and substrate of the membrane, while the dynamic layer over the membrane surface was mainly composed of the retained species, which was the main contributor to the rapid increase in transmembrane pressure in constant permeate flux ultrafiltration. A number of studies demonstrate that the formation of protein layers on the upper surface of membrane is due to the deposition of aggregated and/or

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denatured proteins present in the bulk protein solution. Meireles et al. (1991) studied ultrafiltration of BSA through PS membranes with different MWCOs under different operating conditions. They found that protein denaturation or aggregation and polymerisation were the result of a combination of a variety of parameters, amongst which temperature, crossflow velocity, concentration and time showed significant effects. In the absence of denaturation of protein, no flux decline was observed after the first few minutes. They concluded that long-term fouling in ultrafiltration of albumin was tightly related to the denaturation of protein. Kelly et al. (1993), Kelly and Zydney (1994, 1995) showed that initial fouling in BSA filtration was caused by convective deposition of protein aggregates onto the membrane surface, while native (nonaggregated) BSA only fouled the membrane by chemical attachment to an existing protein deposit via the formation of intermolecular disulfide linkages. Moreover, stirring the BSA solution at an elevated temperature (33 ëC) caused an increase in the number of protein aggregates and significantly increased the protein fouling rate, while prefiltered BSA solution exhibited almost no flux decline after a 30 min filtration (Kelly et al., 1993). GuÈell and Davis (1996) reported the fouling behaviour of BSA, lysozyme, and ovalbumin and their mixtures during microfiltration with different membranes. They found that fouling mechanisms varied with the proteins and membranes used; and protein fouling during microfiltration was not related to the size of the individual protein molecules, but instead to the aggregates that each protein formed under the experimental conditions. In microfiltration of milk, SEM images clearly showed the deposits of protein particles and agglomerates on the membrane surface, at the pore openings and within the membrane pores (James et al., 2003). Numerous studies indicate that the hydraulic resistance of protein deposit can be significantly affected by solution pH and ionic strength. Generally, the resistance is maximal at the protein isoelectric point due to the increase in protein packing density as protein net charge tends to zero. Similarly, when the solution pH is away from the protein isoelectric point, the resistance increases with increasing salt concentration due to the increase in electrostatic shielding between the charged proteins (Jones and O'Melia, 2001; Higuchi et al., 2001; Ricq et al., 1999). In summary, there are two basic mechanisms governing the fouling process, one is internal fouling caused by protein adsorption, deposition at the membrane pore openings and/or within the pores, the other external fouling arising from cake formation over the membrane surface largely due to the aggregates. Both mechanisms may occur simultaneously, and in some cases/stages, one mechanism may dominate, but transition from internal fouling to external fouling may also occur, particularly for high feed concentration and high transmembrane pressure. All these depend on membrane properties, protein properties, solution environment and operating conditions.

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18.4.2 Microbial fouling Microbial fouling can be a major problem in many ultrafiltration and microfiltration systems. Microbial fouling, in particular biofilm formation, can be depicted as a sequence of events, split into a number of steps (Busscher and van der Mei, 2000). Firstly, a conditioning film adsorbs onto the substratum; then bacteria in the liquid are conveyed to the substratum by a number of mechanisms, resulting in non-specific, reversible and eventually specific, irreversible adhesion. In addition to cell-substratum interactions, cell-cell interactions will also occur; and under appropriate conditions, attached microorganisms multiply. Bacteria comprising surface biofilms typically produce extracellular polymeric substances (EPSs), including heteropolysaccharides, lipoproteins, glycoproteins, which serve to anchor cells firmly to the substratum, and further condition the membrane surface so as to stimulate additional microbial colonisation (Costerton et al., 1994). The subsequent colonisation of the membrane surface can occur over a several-weeks to severalmonths period, largely depending on the nutrient status of the feed. Eventually, the growth of the biofilm is controlled by the fluid shear stresses acting on the upper surface of the biofilm (Flemming et al., 1994). Using hydrophobic and hydrophilic strains of known membrane fouling bacteria, Knoell et al. (1999) examined the relationships between bacterial attachment (biofouling potential), water transport, and the surface properties of nine modified polysulfone (MPS) membranes comprising blends of polysulfone (PS) with a sulfonated polyether-ethersulfone/polyethersulfone block copolymer. It was found that hydrophobic mycobacteria attached best to the MPS membranes, but the attachment of both organisms was inversely correlated with the mean aspect ratio of pores, suggesting that irregular or elliptic pores discouraged attachment. Their results indicated that the water flux and biofouling potential of microporous MPS membranes could be manipulated and optimised by modification of the pore geometry and other membrane surface properties. The biofilm formed on membrane surface can significantly affect both the solute and solvent transport through the fouled membrane. Hodgson et al. (1993) clearly demonstrated that biofilms formed of different bacterial types had very different hydraulic resistances, and deposition of just one layer of cells on membrane could result in fluxes as low as 5 l/m2 hr and protein (BSA) rejections as high as 70%. They also showed that the retention characteristics of this bacterial layer could be dramatically reduced upon the addition of EDTA due to modification of the extraculluar matrix. In a review paper prepared by Flemming et al. (1997), it was suggested that the biofilm layer could be conditioned by selecting appropriate surfactant-based cleaners to obtain significant increase in permeate flux, even though the layer thickness was not changed, moreover, increasing shear force may remove the biomass. Any attempts to kill the

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microorganisms attached to the membrane surface should be avoided since microorganisms subjected to low levels of biocides often exude large amounts of extracellular polysaccharides as a protection, and these EPS materials play a key role in promoting cell adhesion and cohesion and constitute the major components of biofilm (Flemming et al., 1997; Baker and Dudley, 1998). In high-density mammalian cell culture, membrane based technologies such as crossflow microfiltration, hollow fibre bioreactor, etc., show great potential in high throughput recombinant therapeutic agent production. However, membrane fouling resulting from interaction between the membrane and the liquid medium hosting cell particles can result in permeate flux decline, and prolonged retention and accumulation of the protein of interest in the bioreactor, which may adversely affect the product quality (Mercille et al., 1994). At the same time the impossibility of regenerating the membrane filters without stopping the culture often leads to very short production cycles. Therefore, the overall performance of the system can be significantly affected. It has also been found that the surface properties of membranes can affect microbial fouling. Avgerinos et al. (1990) found a polymeric ethylenetetrafluoroethylene (ETFE) screen was superior to a stainless steel mesh screen in terms of fouling behaviour. They suggested the hydrophobic nature of an ETFE screen could significantly minimise the attachment of cell to the screen, and thus greatly extend the life of the spin filter. Esclade et al. (1991) also found that in mouse-mouse hybridoma cell culture, stainless steel screens could be more rapidly fouled than polyamide screens. They attributed this difference to surface charge properties as well as their surface hydrophilicity, metals being highly positively charged, polyamides being more neutral and cells being rather negatively charged. There is enormous evidence for hydrophilic surface modification favouring anti-adsorption of protein and anti-adhesion of cells. For example, in high-density perfusion cultures of animal-cells using a stirredtank bioreactor coupled to a hollow fibre microfiltration cartridge, Zhang et al. (1993) found that hydrophilic surface modification by coating the hollow fibre with polyethylene glycol solution prior to autoclaving could reduce fouling rate and prolong the perfusion cultivation time. These differences may be due to the difference between the cell strains they used. Favre (1993) studied the fouling mechanisms of spinfilter in suspended mammalian cell reactors, and found that classical filtration law did not hold for suspended hybridomas in spinfilter filtration. He suggested that fouling was caused by thin cake build up, essentially composed of hybridoma cells and debris; gradual coverage of the filter surface led to complete fouling at a higher TMP. Deo et al. (1996) examined the fouling behaviour of a spinfilter in monoclonal antibody production using an IgM-producing hybridoma and two chimeric IgG-producing myelomas. They showed that single cell suspensions fouled a spinfilter screen, partially but irreversibly, in the early stages of a bioreactor run. Once initial partial screen fouling occurred, further fouling

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continued, although slowly, due to cell growth on the screen surface, and this gradually led to complete coverage. They suggested that the rate of this gradual fouling phase probably depends upon the number of cells deposited on the screen surface during initial fouling. A high perfusion rate and cell density led to high fouling tendency, while an increase in rotational velocity (shear rate) reduced screen fouling. Biological analysis and SEM images revealed that the surface deposits (fouling layer) in mammalian cell culture were composed of DNA or RNA, a lot of dead cells and cell debris, and some living cells (Esclade et al., 1991). The significance of DNA in membrane fouling was also confirmed by Mercille et al. (1994). They demonstrated that the majority of dead cells and some living cells were trapped in DNA-containing aggregates present in protein-free culture, and it was these aggregates that led to the high IgM retention and rapid fouling. Disruption of these aggregates or preventing their formation by the addition of DNase I could significantly reduce membrane fouling. Unlike protein fouling, microbial fouling in mammalian cell culture has not been extensively investigated yet. Given the high shear sensitivity of this type of cell and strict culture conditions, much more work is required to provide a sound basis for promoting membrane application in this field.

18.5 Fouling control Although it seems impossible to avoid fouling in membrane systems completely, many different strategies have been developed to control or lessen fouling. These include the selection of proper membrane, the modification of membrane surfaces, the optimisation of operation conditions, and adjustment and pretreatment of the feed solutions. A fouling layer can also be removed from the membrane surface by chemical cleaning or mechanical methods.

18.5.1 Selection of membrane As stated earlier, membrane material and morphology play a key role in protein adsorption and membrane fouling. Therefore, selection of the suitable membrane is the first step towards the success of a membrane system. In blood purification therapies such as haemodialysis, haemofiltration, membranes with higher permeability and lower fouling tendency are required. It has been reported that a cellulose acetate (CA) membrane induces serious protein adsorption and clot formation when it comes into contact with blood, just like many other membranes. As a result, a decrease in both flux and solute permeability is observed, and infusion of an anticoagulant into the patient during blood purification therapy is required, which leads to serious complications for chronic renal failure patients who require continuous and lengthy treatment (Tsuruta et al., 1993; Deppisch et al., 1998). Whilst membranes made from a methyacrylate

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monomer having a phospholipids polar group, 2-methacryloyloxy-ethyl phosphorylcholine (MPC) were used, protein adsorption, platelet adhesion and their activation were effectively suppressed. These membranes also showed excellent biocompatibility and haemocompatibility (Ishihara et al., 1992; Iwasaki et al., 1997). The effect of membrane materials and properties on fouling has also been demonstrated in the filtration of BSA solutions and cells with different membranes. Nystrom et al. (1998) studied the fouling behaviour of ten different UF and MF membranes in the filtration of a 0.1 g/l BSA solution, and different fouling rates were observed, which were also a strong function of solution pH. Ho and Zydney (1999a) reviewed the work of (i) Bowen and Gan (1991) for flux decline during constant pressure filtration of bovine serum albumin (BSA) solutions through 0.2 m pore size polycarbonate, anodised aluminum oxide (Anopore), and polyvinylidene fluoride (PVDF) membranes; (ii) Davis and his coworkers (MuÈller and Davis, 1996; GuÈell and Davis, 1996) for the fouling characteristics of 0.2 m pore size polycarbonate, polysulfone, cellulose acetate, and PVDF membranes during constant pressure microfiltration of BSA, lysozyme and ovalbumin and their ternary mixtures; and (iii) Hlavacek and Bouchet (1993) for the effects of membrane properties on protein fouling at constant flux. They confirmed that membrane fouling is dependent on membrane material and membrane pore morphology. In pilot-scale operation for clarification of rough beer and pasteurisation of clarified beer using crossflow microfiltration, Fillaudeau and Carrere (2002) studied fouling and retention phenomena during the microfiltration of fermentation broth using ceramic membranes with different pore sizes (0.10, 0.45, 0.80 and 1.40 m). They found that the fouling mechanism differed according to the mean membrane pore diameter. With the 1.4 m membrane, yeast cake resistance was the predominant fouling mechanism and could be decreased by increasing crossflow velocity. With the membranes of pore diameter less than 1 m, the deposition or adsorption of rough beer compounds such as proteins, polyphenols and carbohydrates was the predominant fouling mechanism.

18.5.2 Antifouling modification of membrane surfaces An effective approach against membrane fouling is to change the surface properties of a membrane physically and/or chemically without affecting its transport properties significantly. This can be achieved (Wilbert, 1997) by: (i) introducing charged groups; (ii) increasing hydrophilicity; (iii) introducing steric hindrance; and (iv) biomimetic modifications. In practice, these methods can be utilised alone or in combination. When the surfaces of commercial membranes are modified, a basic principle is that the desired bulk properties, including pore sizes, structures and distribution should be retained.

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Introducing charged surface groups The effects of the charge properties of membranes on membrane fouling have been extensively studied, and have been regarded as an important factor in both fouling control and separation. Generally, operation with a membrane of similar charge to the solute (e.g., protein) can enhance permeate flux and decrease fouling (Marshall et al., 1993). One simple technique for surface charge modification is surfactant adsorption. Chen et al. (1992) pretreated a UF membrane with anionic surfactants, and suggested that small anionic surfactant (AOT) could reduce protein deposition by altering electrostatic interactions between the protein and membrane surface. When used in conjunction with nonionic surfactants or when polyethylene oxide segments were added to their backbone, the anionic surfactants showed significant improvement in flux and fouling resistance compared with that of the single AOT or the non-ionic surfactant. They also found that effectiveness of pretreatment was sensitive to pH and protein charge. Charged groups can also be introduced by chemical modification. Nakao et al. (1988) prepared a negatively charged ultrafiltration membrane from sulfonated polysulfone (SPS) and a positively charged ultrafiltration membrane from polysulfone having quaternary ammonium groups (APS). Experimental results showed that in the ultrafiltration of a BSA solution fouling was low at pH 10.3 and 3.8 for SPS and APS membranes, respectively, due to an electrostatic repulsion effect between the membrane and the protein. Higuchi et al. (1990) chemically modified both inner and outer surfaces of polysulfone hollow fibres with propane sulpone to introduce negatively charged -CH2CH2CH2SO3ÿ segments on the membrane surfaces. Experimental results indicated that the modified fibres showed better performance for anti-absorption of protein than the unmodified fibres. Increasing hydrophilicity Hydrophilic membranes generally show less fouling and higher flux (Gekas and HallstroÈm, 1990; Babu and Gaikar, 2001). Therefore, hydrophobic membranes may need to be modified by introducing hydrophilic segments at their surface and/or within their matrices, whereas the hydrophilic modification only at the membrane surface is preferred because it is easier, inexpensive and scalable, and the mechanical strength of the membrane will normally be unaffected. The hydrophilic modification of a membrane can be achieved by using coating, blending and grafting techniques. Physical coating This is the simplest method for modifying surface properties of substrates. It can be achieved by permeating through the membrane the solution of hydrophilic

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materials such as hydrophilic polymers (Kim et al., 1988; Hanemaaijer et al., 1989; Brink et al., 1993; Brink and Romijn, 1990), surfactants (Brink and Romijn, 1990; JoÈnsson and JoÈnsson, 1991; Chen et al., 1992), or proteins (Ghosh and Cui, 1998), etc. In the filtration of a whey-protein solution using a PS ultrafiltration membrane, Brink and Romijn (1990) found that the membrane pretreated with non-ionic, hydrophilic polymers could minimise protein adsorption and decrease the membrane resistance during ultrafiltration, while the application of surfactants and ionic polymers was generally less successful. Reddy et al. (2003) modified polyethersulfone ultrafiltration membranes by preadsorption of poly(sodium 4-styrenesulfonate) (PSS) upon the permeation of aqueous solution of the polymer for about 100 min, and found that the surface modified membranes had excellent `cleanability' and antifouling characteristics compared to unmodified membranes. However, such an approach to modifying surface properties has generally been found inapplicable in practice because the resulting coating tends to be temporary, and is removed as a whole or in part after initial use, as shown by Kim et al. (1988). Blending The surface hydrophilicity of membranes can also be improved by blending a hydrophilic polymer into the bulk polymers when fabricating the membranes. The advantage of this approach is that the pore size and its distribution of membranes can be controlled as required as in the normal membrane fabrication process, and the hydrophilic component can be distributed evenly both on the membrane surface and within the matrix. Polyvinyl-pyrrolidone (PVP) is a commonly used hydrophilic polymer for preparing blended membranes (Lafreniere et al., 1987; Wienk et al., 1995; Marchese et al., 2003). Qin et al. (2003) prepared polyethersulfone (PES)-PVP hollow fibre membranes by spinning blended PES-PVP solution and then posttreated the membranes with a hypochlorite solution over a range of concentration to remove the PVP for a fixed period of 48 h. Experimental results for BSA filtration indicated that untreated membrane showed very low fouling tendency due to the hydrophilic nature of the PVP/PES blend, but the treated membrane where PVP was removed experienced a significant fouling tendency. However, like physical coating, blended membranes are subject to the elution of the hydrophilic component, leading to increase in fouling rate. As Kobayashi and Tanaka (1995) pointed out, the elution of hydrophilic polymer causes problems and could even be dangerous for clinical applications. Therefore, it is desirable to avoid the elution of the hydrophilic component from the modified membranes. Surface grafting Compared with other approaches, chemical modification by surface grafting to introduce hydrophilic segments onto the surface of hydrophobic membranes

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exhibits the advantages of both the hydrophilic surface and a strong hydrophobic membrane (Higuchi and Nakagawa, 1990; Higuchi et al., 1992). The original characteristics of mechanical strength and thermal stability are retained, since only the membrane surface is modified (Higuchi et al., 1988). In addition, the introduced hydrophilic segments are more stable and not easy to elute, since they are chemically bonded on the surface. Surface grafting has hence become a popular approach to modify the surface properties of membranes. Using graft polymerisation, a variety of hydrophilic monomers have been grafted to membrane surfaces to increase their hydrophilicity and reduce their potential to foul during filtration. Monomers employed include: · acrylic acid (AA) (Wavhal and Fisher, 2002; Xu et al., 2002; Freger et al., 2002; Ulbricht et al., 1998); · methacrylic acid (MA) (Freger et al., 2002; Ulbricht and Belfort, 1996; Yamagishi et al., 1995a); · vinyl acetate (Kim et al., 1991); · glycidal methacrylate (GMA) (Kim et al., 1991; Yamagishi et al., 1995a,b); · 2-hydroxyethyl methacrylate (HEMA) (Ulbricht et al., 1996; Wang et al., 2000); · poly(ethylene glycol) methacrylate (PEG-MA) (Ulbricht et al., 1996; Gilron et al., 2001); · 3-sulfopropyl methacrylate (SPM) (Wang et al., 2000; Gilron et al., 2001; Freger et al., 2002); · 2-dimethylamino-ethyl methacrylate (AEMA) (Ulbricht et al., 1998); · 2-trimethylammonium-ethyl methacrylate chloride (AmEMA) (Ulbricht et al., 1998); · 2-acrylamido-2-methyl-1-propanesulfonic acid (AMPS) (Gilron et al., 2001); and · N-vinyl-2-pyrrolidinone (NVP) (Pieracci et al., 1999, 2002). Grafting occurs at active ionic or radical sites at the membrane surface. Generation of such sites may be accomplished by several different approaches, including: · UV radiation (Yamagishi et al., 1995a,b; Ulbricht et al., 1996, 1998; Pieracci et al., 1999, 2002); · oxidants and/or chemical redox systems such as ozone (Wang et al., 2000; Xu et al., 2002); · low temperature helium (Ulbricht and Belfort, 1996) or argon (Wavhal and Fisher, 2002) plasma; · ionising radiation (Kim et al., 1991; Kabanov and Kudrayavtsev, 2003). Photo-induced grafting by UV radiation is a particularly useful technique for the surface modification of membrane due to its low cost of operation and mild reaction conditions. Another advantage of UV radiation is the ability to control

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trunk polymer bond cleavage by judicious selection of both emission intensity and wavelength (Pieracci et al., 2002). Pieracci et al. (2002) prepared low protein fouling 50 kDa PES ultrafiltration membranes by UV-assisted graft polymerisation of the monomer NVP using a dip modification technique. The effects of wavelength and filters were also discussed. Kaeselev et al. (2001) modified PES and PS ultrafiltration membranes by UV-assisted graft polymerisation of three different hydrophilic monomers, NVP, AMPS and 2acrylamidoglycolic acid monohydrate. Four different modification conditions were identified to produce modified UF membranes with filtration performance superior to the base PES, the base PS, or regenerated cellulose (RC) control membranes in terms of their low fouling and excellent cleaning characteristics. Higuchi and his coworkers (Higuchi and Nakagawa, 1990; Higuchi et al., 1988, 1990, 1992) conducted a series of work to introduce hydrophilic segments chemically on both the inner and outer surfaces of PS hollow fibres to improve the protein resistance. They (Higuchi et al., 2002) modified PS hollow fibres by chemically bonding PVP on the surface and found PVP/PS membranes to be the most hydrophilic among the PS and surface-modified PS membranes prepared in their study. The PVP/PS membranes not only exhibited lower protein adsorption from a plasma solution, but also showed a more suppressed number of adhering platelets on the surface than PS and other surface-modified membranes. Using ozone treatment to introduce peroxide onto the membrane surface, HEMA could be grafted onto polypropylene (PP) flat sheet MF membranes (Wang et al., 2000). The graft was initiated at a mild temperature by redox decomposition of peroxide. The HEMA grafting made the surface of the PP membrane hydrophilic and anti-adsorption of protein (BSA); these effects were dependent on the ozone-treating duration. Modification of polymer surfaces can be rapidly and cleanly achieved by plasma treatment due to the possibility of forming various active species on the surface of polymers. Gancarz et al. (1999) modified polysulfone membranes with acrylic acid (AA) using three different plasma-initiated graft polymerisation methods: (i) grafting in solution; (ii) grafting in vapour phase; and (iii) plasma polymerisation. They demonstrated that a polysulfone membrane with AA plasma-initiated grafting in a vapor phase of monomer seemed to be the most promising from the point of view of filtration properties. Wavhal and Fisher (2002) obtained a complete and permanent hydrophilic modification of PES membranes by argon plasma treatment followed by polyacrylic acid (PAA) grafting in the vapour phase. The modified membranes were less susceptible to protein fouling than unmodified membranes, and pure water flux for the modified membranes was markedly increased. Furthermore, modified membranes were easier to clean and required little caustic to recover permeation flux. They also suggested that vapour phase grafting yields were much higher than solution phase grafting yields.

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Introducing steric hindrance Due to steric hindrance effects, hydrophilic polymers grafted onto the surface of polymer substrates can considerably reduce the adsorption of proteins and adhesion of cells (Kishida et al., 1992). The conformation of graft chains in a good solvent can dramatically change with graft density. With increasing density of hydrophilic polymers grafted onto a hydrophobic surface, graft chains are obliged to stretch away from the surface, forming a hydrophilic polymer `brush' on the surface (Milner, 1991). In this so-called `brush regime', a high degree of protein rejection is generally observed for a variety of proteins (Stengaard, 1988; Bearinger et al., 1997; Currie et al., 1999). Polymer brushes on solid surfaces can be prepared by either by reversible adsorption (physical sorption) of diblock or triblock copolymer chains on the surface or chemical grafting (Zhao and Brittain, 2000). Generally, polymer brushes prepared by adsorption exhibit thermal and solvent instability due to weak interactions between the substrate and the block copolymers, while polymer brushes prepared by chemical grafting are covalently tethered to the membrane, thus the polymer brushes produced are very stable. Hydrophilic brushes such as PEO attached to a hydrophobic substrate can give the surface an extraordinary ability to resist protein adsorption (Wang et al., 2002; Kim and Kim, 2002; Jeon et al., 1991) and cell adhesion (Nakayama et al., 1999). Extensive studies have been devoted to create polymer surfaces that tether PEO chains. Lee et al. (1989) found that pendant PEO chains on the material's surface played an important role in reducing blood proteins adsorption on the surface. This property is believed to have resulted from the combination of hydrophilicity, steric repulsion between the grafted hydrophilic polymer brushes and proteins, and a unique coordination with surrounding water molecules in an aqueous medium (Lee et al., 1995). Using a photo-induced grafting technique, Ivanchenko et al. (1995) prepared PU films grafted with poly(methoxy-PEG methacrylate) chains with 4, 9, and 23 units of ethylene glycol, and found that protein adsorption and platelet adhesion to the grafted surfaces were reduced mostly when the graft yield was largely reduced by the use of a high concentration of a chain transfer agent. In addition, extraordinarily high graft yields were not effective in preventing protein adsorption, which was attributed to the migration of proteins into the graft layer. Interestingly, membrane permeability can be actively regulated by the dynamic transition of graft chains. By grafting PAA onto a porous Nucleopore membrane, Iwata et al. (1998) studied the effects of graft density and pH on the filtration performance of the modified membrane. They found that PAA chains grafted on the membrane surface dynamically changed their configuration in response to medium pH. AFM images demonstrated that graft chains shrank and precipitated on the surface of the membrane and the wall of pores at acidic pH, thereby opening the pores of the membrane, whereas they hydrated and thus

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18.4 Schematic representation of a molecular valve capable of regulating the membrane permeability and rejection (adapted from Iwata et al., 1998).

effectively closed the pores at neutral and alkaline pH. As shown in Fig. 18.4, the PAA graft chains dynamically opened and closed the pores in response to the change in medium pH, functioning as a molecular valve to regulate the permeation characteristics. Biomimetic modifications One promising and exciting development in biomaterials is to mimic a biological surface in nature. This is particularly useful not only for artificial organs, but also for medical devices in biomedical applications such as haemodialysis and plasmapheresis. For example, the red blood cell plasma membrane, unlike synthetic polymer membranes, naturally resists protein fouling. This property is attributed to the unique phospholipid bilayer structure of the cell membrane (Hayward and Chapman, 1984). Inspired by the surface structure of biomembranes, a series of new biomaterials with excellent properties of protein adsorption and cell adhesion resistance have been developed by Ishihara and his coworkers (Ishihara et al., 1990; Iwasaki et al., 1997), based on the mimicking of a simple component present on the extracellular surfaces of the lipid bilayer that forms the matrix of the plasma membranes of cells, namely, the phosphorylcholine group of phosphatidylcholine and sphingomyelin. Ishihara et al. (1990) copolymerised 2methacryloyloxyethyl phosphorylcholine (MPC) with n-butyl methacrylate (BMA) and prepared a hydrogel membrane with this poly(MPC-co-BMA) (PMB) by a solvent evaporation method. This hydrogel membrane allowed water soluble organic compounds and proteins of molecular weights less than 10 kDa to pass through, whilst the protein with molecular weight of more than 100 kDa would be retained. A detailed description about lipid bilayer structure of biomembrane, the functions and applications of polymerised phospholipids has been presented by Chapmann (1993), showing that phosphorylcholine treatments can be successfully applied to a wide range of materials for medical uses, such as intravascular catheters and filtration membranes.

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In order to improve their surface blood compatibility and protein adsorption resistance, a number of conventional polymeric materials used in the biomedical field were modified by blending with the MPC polymers, these include MPC modified PS membranes (Ishihara et al., 1999a,b; Hasegawa et al., 2001), MPC modified polysulfone (PS) hollow fibre membranes (Ishihara et al., 2002), and MPC modified cellulose (CA) membranes (Ye et al., 2003). The mechanical strength of these blended polymer membranes has not been changed, while the MPC polymer can serve as a doubly functional polymeric additive, i.e., to generate a protein-adsorption-resistant characteristic, and to increase the membrane hydrophilicity. All these MPC modified membranes could significantly decrease the protein adsorption at the surface of the membranes, and platelet adhesion and activation can also be effectively suppressed. As a result, the permeate flux and protein transmission could be maintained almost the same during the observed period of ultrafiltration of a protein mixture (Ye et al., 2003). Ye et al. (2003) examined the amount of protein and platelet adsorption on three different membranes, CA I, CA/PMB30 III and polysulfone membranes. CA I was made from 20% cellulose acetate only, while CA/PM30 III was made from 19% CA and 1% poly(2-methacryloyloxyethyl phosphorylcholine (MPC)-co-n-butyl methacrylate (BMA)) (PMB30). They found that the amount of each protein absorbed on the CA/PMB30 membrane surface was significantly less than those absorbed on the CA and polysulfone membranes for all of three proteins examined, i.e., the amounts of absorbed albumin on CA/ PMB30 III, CA and PS membranes were around 0.5, 0.9 and 2.4 g/cm2, respectively; for absorbed -globulin, the amounts were around 0.1, 1.5 and 1.9 g/cm2, respectively; for absorbed fibrinogen, the amounts were around 0.3, 1.1 and 0.9 g/cm2, respectively. Their results suggested that the CA/PMB30 blend membrane had little interaction with blood components such as protein and platelets, thus showing good haemocompatibility in terms of suppression of protein adsorption, platelet adhesion and its activation. Similar results were also reported for MPC modified polysulfone membranes (Ishihara et al., 1999b). In addition to blending, MPC polymer can also be coated on other substrates (Chapmann, 1993). Dudley et al. (1993) treated PVDF with a surface grafting of phosphorylcholine. The experimental results indicated that membranes with this non-protein-binding coating showed less severe flux declines and a 95% reduction in protein adsorption. Reuben et al. (1995) modified microfiltration membranes of cellulose triacetate, PS and PVDF by surface grafting MPC after plasma etching with oxygen. The performances of these modified membranes were evaluated with water, buffer, bovine serum albumin (BSA), yeast fermentation broth, beer and orange juice. They found MPC modified PVDF membranes gave the best results. In most cases, the MPC coating PVDF showed both higher initial flux and lower fouling rate. Less protein adsorption and deposition than untreated membranes were shown by electron microscopy. Akhtar et al. (1995) also treated PVDF and CA membranes by surface grafting

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MPC onto those membranes. The experimental results for the filtration of BSA solutions indicated that phospholipid coating improved the flux decline of both PVDF and cellulose acetate membranes. They also observed that this coating was more effective on the PVDF membrane. Moreover, a reduced protein fouling at the surface and within the matrix of the coated membranes was demonstrated as assessed by anionic gold staining. Therefore, the phospholipid coating could be an effective treatment for both reducing protein fouling and improving their performance in filtration. Other examples of biomimetic modifications of membranes include the introduction of heparin (HE), the most widely used blood anticoagulant, or endothelial cell surface heparan sulfate (ESHS) to prepare HE or ESHS containing membranes (Jen et al., 1998; Baumann and Kokott, 2000), and the preparation of membrane consisting of biologically derived matrices (Attafuah and Hall, 1995) to mimic the functions of normal blood vessels. Baumann and Kokott (2000) prepared HE and ESHS containing polysulfone, polycarbonate and polyurethane (PU) membranes, and found that the HE and ESHS modified membranes showed much less protein and platelet adhesion, and the ESHS coating membrane may be used without anticoagulants. In order to mimic the anti-coagulation property of blood vessels, UF membranes composed of biologically derived matrices from a bovine aorta endothelial cell line were also prepared, and their performances were demonstrated to be superior to CA ultrafiltration membranes in terms of fouling and rejection properties (Attafuah and Hall, 1995).

18.5.3 Optimisation of operating conditions As fouling is significantly affected by operating conditions, it is necessary to optimise these operating parameters to minimise fouling during processing operations. Parameters needing to be optimised include feed concentration, solution environment (i.e., buffer selection, pH, ionic strength) and transmembrane pressure (permeate flux) and/or crossflow velocity. Since fluid dynamics play a key role in membrane fouling control, enormous efforts have been devoted to hydrodynamic management in membrane filtration. This topic will be separately addressed in the next section. In biotechnological applications, membrane separation aims at cost-effective production, this is generally evaluated in terms of productivity (yield) and product quality (purification factor). However, in most cases, fouling control has been mainly focused on minimising flux decline, although higher flux does not necessarily give higher productivity and good-quality product. This is true for bioproducts separation from fermentation broths, particularly for protein fractionation using ultrafiltration, since both protein transmission and separation selectivity are flux dependent (Wan et al., 2002; Ghosh et al., 2003). Therefore, optimisation of operating conditions is not straightforward and normally

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involves extensive experimentation, which is both time consuming and expensive due to relatively large amounts of protein consumption. Two new experimental techniques have been developed for assessing membrane suitability and optimising operation parameters. One is pulsed sample injection ultrafiltration (Ghosh and Cui, 2000b), which allows a detailed analysis of fouling and protein transmission at the cost of a small amount protein consumption. The other is carrier phase ultrafiltration (CPUF) (Ghosh et al., 2003), which enables rapid parameter (pH and ionic strength) scanning experimentation to determine the optimal conditions. Both of these are based on constant permeate flux ultrafiltration. The usefulness and effectiveness of the combination of pulsed sample injection technique with CPUF have been demonstrated in the fractionation of human serum albumin (HSA) and human immunoglobulins (HIgG) (Wan et al., 2002; Ghosh et al., 2003), bovine serum albumin (BSA) and monoclonal antibody Campath-1H (Ghosh et al., 2003). High-performance tangential flow filtration (HPTFF) is another development of conventional tangential flow filtration (TFF) (also referred to as crossflow filtration) (van Reis et al., 1997). However, unlike conventional TFF, which typically operates in the pressure-independent part of the permeate flux curve, HPTFF is carried out in the pressure-dependent flux regime. In the HPTFF system, the formation of a pressure gradient along the length of feed channel is overcome by operating a co-current permeate stream that maintains transmembrane pressure constantly along the length of the TFF module. Thus both high selectivity and high mass throughput can be obtained since the permeate flux and the local transmembrane pressure can be carefully controlled at optimum levels. While membrane fouling is controlled by selecting appropriate fluid dynamic and mass transfer conditions. A HPTFF process development includes membrane selection, buffer optimisation, fluid dynamic optimisation and the economic evaluation. By exploiting both size and charge mechanisms, HPTFF can be used to separate monomers from dimers based solely on size differences and proteins of equal size on the basis of charge differences (van Reis and Zydney, 2001).

18.5.4 Flow manipulation The role of hydrodynamics in membrane systems is crucially important in controlling concentration polarisation and fouling (Belfort, 1989; Al-Bastaki and Abbas, 2001; Cui et al., 2003). Concentration polarisation has been shown to have a significant effect on permeate flux and fouling in microfiltration and ultrafiltration systems. Higher concentration polarisation generally results in decreasing of permeate flux and increasing of fouling potential. As in heat transfer, any technique which interrupts the formation of a continuous boundary layer on the transfer surface and enhances mixing of the fluid across the flow channel, is likely to reduce concentration polarisation and to enhance permeate

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flux. So far, various techniques for improvement of system hydrodynamics have been investigated. These include vortex mixing (Balakrishnan and Agarwal, 1996a,b; Costigan et al., 2002), tube inserts (Najarian and Bellhouse, 1996a,b; Bellhouse et al., 2001), pulsatile flow (Wang et al., 1994), and gas sparging (Cui and Wright, 1994, 1996; Mercier et al., 1997; Ghosh et al., 1998). Most of these techniques have been proved effective in both enhancing permeate flux and reducing fouling in membrane filtration. Increasing the stirring speed in a stirred cell ultrafiltration unit is the simplest way to enhance mixing. Slater et al. (1986) found that in the fractionation of bovine alkaline phosphate (MW 140000) and BSA in a stirred cell UF module with a 100 kDa MWCO regenerated cellulose membrane, the efficiency of fractionation was higher at higher speeds. In the purification of lysozyme from chicken egg white (CEW) using ultrafiltration, permeate flux was found to increase with an increase in stirring speed; while the transmission of lysozyme was found to decrease with the increase in stirring speed (Ghosh and Cui, 2000c), which was explained in terms of the concentration polarisation effect for this protein. Vogel and and Kroner (1999) showed that in the culture of mammalian cell with controlled shear filtration, permeate flux increased with an increase in rotor speed. Based on batch harvesting results, no evidence of fouling was found and the product transmission was 100% by selecting appropriate conditions including the rotor speed. In dynamic filtration of polysaccharides using microfiltration and ultrafiltration membranes, Brou et al. (2003) demonstrated that both permeate flux and solute flux could be significantly increased with an increase in rotating speed. Balakrishnan and Agarwal (1996a) adopted a vortex flow filtration device fitted with a 100 kDa MWCO polyacrylonitrile membrane to investigate the effects of system hydrodynamics on permeate flux and protein transmission in single component ultrafiltration systems. Taylor vortices were generated in their experiments to affect the system hydrodynamics. Lower protein transmission was reported at higher rotation speeds and higher axial velocities. The effect of permeate flux on transmission was also found to be significant. It was explained by a combined concentration polarisation ± irreversible thermodynamics model. In a subsequent paper (Balakrishnan and Agawal, 1996b), they investigated the fractionation of simulated mixtures of lysozyme/ovalbumin and lysozyme/ myoglobin utilizing Taylor vortices, and found that the ultrafiltration characteristics of dilute protein mixtures were virtually identical to those of their individual components at low transmembrane pressures and high membrane rotation speeds. The selectivity of separation was controlled primarily by the extent of polarisation of the smaller, preferentially transported species (lysozyme). Gehlert et al. (1998) examined the effect of Dean vortices produced by flowing through curved tubes on ultrafiltration and microfiltration of polysaccharides, proteins and yeast suspensions. Their results showed that

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helical hollow fibre membrane modules using Dean vortices could effectively reduce concentration polarisation and fouling. Flux enhancements ranging from 20 to 200% for dextran T500, up to 85% for BSA solution and 30 to 90% for bear yeast (MF only) were observed. Kaur and Agarwal (2002) studied the effect of Dean vortices on protein transmission in a thin channel flow module. The model protein used was lysozyme and the membrane was a hydrophilic cellulose acetate membrane with a molecular weight cutoff of 30 kDa. Experimental results indicated that, as in the case of Taylor vortices, Dean vortices could be effective in minimising concentration polarisation and enhancing mass transfer. Bellhouse and co-workers (Millward et al., 1995a; Najarian and Bellhouse, 1996a; Bellhouse et al., 2001) carried out a detailed study of the performance of helical screw thread inserts in tubular membranes. The design combined predominantly helical flow, in which Dean vortices were generated. Experimental results showed that a dramatic increase in permeate flux (by factors of 6±10) could be obtained under typical microfiltration, ultrafiltration and nanofiltration conditions. Helical screw thread flow promoters were designed for processing shear-sensitive fluids. They are simple to construct and operate well under laminar, quasi-steady flow conditions. When used for the ultrafiltration of 60 g/l BSA in tubular membrane geometry (Millward et al., 1995a), no long-term fouling was observed and significant permeate flux enhancement was obtained. Millward et al. (1995b) also investigated the enhancement of plasma microfiltration using oscillatory flow and flow deflectors, which were used to generate vortex waves. The vortex waves were found to be effective in improving permeate flux and a flux enhancement factor of 3.5 relative to a flat unobstructed channel was reported. Najarian and Bellhouse (1996b) examined the effects of pressure pulsation and flow deflector on fractionation of bovine plasma and -globulin using a flat-plate ultrafilter equipped with a ladder-like flow deflector. Application of transmembrane pressure pulsation improved the selectivity of albumin-globulins by a factor of 3 with a GR40PP membrane (100 kDa MWCO polysulfone membrane), permeate flux was also increased. Rodgers and co-workers studied ultrafiltration of protein solutions using backpulsing with frequencies up to 5 Hz. Rodgers and Sparks (1991, 1992) studied the effects of transmembrane pressure pulsing on protein ultrafiltration, and found that transmembrane pressure pulsing was very effective in reducing both fouling and concentration polarisation either using fully or partially retentive membranes. Wilharm and Rodgers (1996) examined the effects of varying pulse amplitude and duration in pressure pulsation ultrafiltration of BSA and IgG. It was found that the variation in pulse duration did not significantly affect permeate flux but the mass flux of BSA could be increased by a factor of three. No data was presented for the transmission of IgG. It was suggested that transmembrane pressure pulsing might enhance solute flux by removing lodged solute molecules from the membrane pores.

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Transmembrane pressure pulsing has also been shown to be effective in microfiltration for cell harvest and removal. The term `backflushing' refers to low-frequency permeate flow reversal (normally a few seconds once every few minutes or more), and the term `backpulsing' is used when the frequency is high (less than one second once every few seconds or less) (Redkar and Davis, 1995). Kroner et al. (1984) achieved 50% enhancement in the net flux during the removal of E. coli bacteria from a fermentation broth using crossflow microfiltration with backflushing by reversing the transmembrane pressure for 5 s every 5 min. Matsumoto et al. (1987, 1988) reported up to a tenfold flux increase with backflushing (for 5 s every 3 min) for yeast suspensions. Redkar and Davis (1995) studied backpulsing process for microfiltration of washed yeast suspensions. Under optimum conditions, flux enhancement of approximately 40-fold was observed. In microfiltration of bacterial lysates using backpulsing, more than tenfold increase in the net flux and improved protein transmission from 60 to 100% were obtained with backpulsing under optimum operating conditions (Parnham and Davis, 1996). Kuberkar et al. (1998) examined the adhesive foulants on backpulsing performance. Washed E. coli suspensions and diluted and undiluted E. coli fermentation broths were used. Experimental results showed that under optimum backpulsing conditions, the net fluxes for the washed bacteria were approximately tenfold higher than those obtained during normal crossflow microfiltration operation, whereas only a twofold improvement in the net flux was achieved for the fermentation broths. They attributed it to extracellular macromolecules leading to irreversible fouling and more adhesive cake. Kuberkar and Davis (2001) studied microfiltration of yeast suspensions, BSA solutions, and mixtures of yeast and BSA with or without backflushing or crossflushing. It was found that backflushing was more effective than crossflushing in removing internal and external fouling caused by BSA, yeast or both. In the filtration of BSA-yeast mixture, backflushing could effectively enhance the permeate flux and reduce fouling only at the early stages of operation, while at longer times, the net flux for the experiments with backflushing declined sharply so that it became comparable to (or even lower than) the flux obtained without backflushing. It seems that the efficacy of backflushing or backpulsing could be significantly undermined by the complexity of biological solution. The effect of backflushing or backpulsing in the filtration of real fermentation broths needs to be further examined. The creation of gas-liquid two-phase flow in a membrane module by sparging air or another gas has been shown to be an effective way for controlling concentration polarisation (Cui and Wright, 1994; Bellara et al., 1996; Mercier et al., 1997). Bellara et al. (1996) employed gas sparging in crossflow hollow fibre ultrafiltration of dextran and albumin solutions. Flux enhancements of 20± 50% for dextran and 10±60% for albumin were obtained. In a downward flow of feed and gas bubbles with a tubular ultrafiltration membrane, a flux increase of

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up to 320% was observed (Cui and Wright, 1994), compared with conventional operation. They found that the low flowrate gas sparging was most effective in enhancing ultrafiltration in the liquid laminar flow region, and pointed out that the two-phase flow pattern was an important parameter in determining the performance of sparged ultrafiltration. In a similar study carried out with organic hollow fibers by Cabassud et al. (1997), it was found that air slugs were efficient in enhancing mass transfer when it was limited by particle deposit. The significant increases in ultrafiltration flux (of about 110%) were supposed to be linked to the high wall shear stresses induced by the air slugs. In fractionation of protein, both permeate flux and separation selectivity can be enhanced with gas sparged ultrafiltration. It was reported that in the fractionation of HSA and IgG (Li et al., 1997), BSA and lysozyme (Ghosh et al., 1998), introduction of gas bubbles greatly increased the selectivity of the fractionation and permeate flux. Under optimal conditions complete separation of these two binary protein mixtures were obtained (Li et al., 1997; Ghosh et al., 1998). The mechanism of enhancement was explained in terms of disruption of the concentration polarisation layer and enhanced mass transfer due to bubbleinduced secondary flow (Ghosh et al., 1998). Lee et al. (1993) studied air slugs entrapped in crossflow filtration of bacterial cell suspensions (E. coli and Brevibacterium flavum). When flat ultrafiltration (MWCO 300 kDa) membranes were used, a twofold enhancement in flux was obtained, while in microfiltration with a membrane of 0.2 mm pore diameter the gain was 1.3. Mercier et al. (1998) studied the use of an upward gas/liquid slug flow to reduce tubular mineral membrane fouling in yeast suspension filtration. They found that when external fouling was the main limiting phenomenon, flux enhancements of a factor of three could be achieved with gas sparging for both ultrafiltration and microfiltration systems, compared with single liquid phase crossflow filtration. In the same study, they demonstrated that when coupled in continuous alcoholic fermentation with cell recycle, sparged ultrafiltration allowed a high and stable flux to be maintained over 100 h of fermentation, with a final cell concentration of 150 g dry weight/l. At the same biomass level, a twofold gain in flux could be obtained at half the cost of conventional single phase crossflow filtration in terms of energy consumption. The introduction of gas/liquid two-phase flow has been shown to enhance significantly the performance of many membrane process applications. A comprehensive review on gas bubbling in membrane systems, including its mechanism, modelling and applications, has been presented by Cui et al. (2003). In addition to the above-mentioned approaches, fouling control can also be realised by other novel module designs such as a tubular design with a `scoop' (Schubert and Todd, 1980) and filters with internal helical grooves (Costigan et al., 2002), and by additional force field, e.g., electrical field (Wakeman and Tarleton, 1986; Mameri et al., 1999).

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18.5.5 Membrane cleaning Fouled membranes need to be cleaned to restore their performance after the flux and/or selectivity drop to some minimally acceptable levels. Cleaning can be accomplished by physically removing the foulants from the membranes (e.g., backflushing, mechanical scrubbing) or by chemical cleaning agents (e.g., acids, alkalies, surfactants, enzymes) to remove or decompose the foulants. Physical cleaning Backflushing uses reversed liquid flow to flush foulants from and out of membrane surfaces. It is simple and generally very effective for removing the cakes of particulates from the surfaces, it can also remove foulants from the membrane interior. Backflushing can be performed during a filtration operation with the permeate or air, or in a separate cleaning cycle with either rinse water or suitable cleaning solutions. Backpulsing (Rodgers and Sparks, 1993; Redkar and Davis, 1995) is a variation of traditional backflushing. This method employs a back pressure of an extremely rapid pulse (pulse duration generally less than 1 s) every 10±30 s throughout the process, thus foulants can be removed periodically by the back flow during operation. It has been proved that this method is very effective for membrane cleaning and flux enhancement, in the microfiltration of washed cells (Redkar and Davis, 1995, Kuberkar et al., 1998) and protein ultrafiltration (Rodgers and Sparks, 1993). In addition to the hydrodynamic methods mentioned above, other physical cleaning methods such as pulsed electric field (Bowen and Sabuni, 1992; Bowen and Ahmad, 1997) and ultrasonic field (Kobayashi et al., 2003) can also be used. Chemical cleaning A variety of chemicals can be used for membrane cleaning, which include acids, alkalis, surfactants, chelating agents, enzymes and some oxidising or reducing agents etc. The most commonly used chemicals for membrane cleaning are HCl, H2SO4, NaOH, EDTA, proteases, NaClO and NaHSO3. Chemical reactions involved in cleaning include hydrolysis, peptisation, saponification, solubilisation, dispersion, and chelation (TraÈgaÊrdh, 1989). The characterisation of foulants is the key to the selection of cleaning solution. Unfortunately, most fouling phenomena are poorly understood and poorly characterised. Generally, cleaners composed of mineral acids, sodium hexametaphosphate, polyacrylates and EDTA are suitable for salt precipitates and mineral scalants while caustics based on NaOH and /or NaClO are used for removing fats and proteins; enzyme cleaners are used to remove proteins and other types of biofoulants in specific instances (Ho and Sirkar, 1992).

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Bartlett et al. (1995) examined caustic cleaning, alkali/acid sequence cleaning and formulated agent cleaning for flux restoration of sintered stainless steel and ceramic microfiltration membranes fouled by whey proteins. Their results suggested that the use of a caustic multicomponent cleaning agent would provide efficient cleaning of inorganic microfiltration membranes under optimised conditions. MunÄoz-Aguado et al. (1996) examined the cleaning of ultrafiltration membranes fouled by BSA and whey with surfactant and enzyme. Their results showed that it was most effective to clean first with an enzyme and then with a detergent, or, if both were present in the same cleaner, appropriate formulation was required to avoid interfering with any others. Chen et al. (2003) applied a statistical factorial design to identify the key factors and their interactions in both physical and chemical cleaning of ultrafiltration (UF) and reverse osmosis (RO) membranes. By using the optimised membrane cleaning methods, the cleaning efficiency was significantly improved, and higher membrane filtration capacity and efficiency were obtained. In industrial applications, membranes are generally cleaned by cleaning-inplace (CIP) procedures to minimise downtime, and cleaning solutions are circulated without pressure to prevent deep penetration of the foulants into the membrane. While in clinical uses (e.g., dialysers), cleaning operations are normally performed off-line for reuse, or more often, the fouled membranes are simply discarded.

18.6 Conclusion and future trends Fouling is a major problem in most membrane-based operation systems, resulting in low operating efficiency and even system failure. Fouling also constitutes a main obstacle to wider applications of membrane technology. Fouling is affected by a variety of factors, including the membrane itself, solution environment, system hydrodynamics and operational flux. Fouling can be controlled to some extent by selecting appropriate membranes, manipulating fluid hydrodynamics and backflushing, optimising solution environment (e.g., pH and ionic strength) and modifying the membrane surface physically or chemically. Unfortunately, chemical cleaning is still inevitable to restore membrane performance, but cleaning increases operational complexity and cost, and may reduce the service life of the membrane. Fouling, though extensively studied, is still poorly understood. Given the complexity of biological solutions and products, more fundamental research is required, particularly the fouling mechanisms involved. Biological solutions and/or fermentation broths are complicated systems containing proteins, EPS, and/or cell and cell debris, etc. The solution chemistry governs the stability and interactions among the constituents. A better understanding of protein properties and their interaction with cells and membranes under different conditions (e.g., pH, ionic strength and shear stress) will provide a scientific basis for fouling

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control in engineering aspects. A better understanding of fouling mechanisms also provides valuable guidance in the selection of cleaning agents and cleaning protocols. New anti-fouling orientated module and process design and development are required. More efficient fouling control also relies on the development of new membrane materials and module designs based on specific applications. Biomimetic modification and polymer brush seem quite promising and need to be further investigated, which will probably facilitate the application of membrane systems in both biological and biomedical fields. The development of cheap and disposable membrane may be an alternative as well. The understanding of how the operating conditions, in particular, permeate flux, affect membrane fouling, has been greatly improved in recent years. The concept of critical flux marked an important step towards this improvement (Field et al., 1995) and now is generally accepted. This concept has led to a complete U-turn in the methodology of membrane system design and operation, from pursuing high flux to operating at low flux to eliminate membrane cleaning. This practice is now widely accepted in membrane bioreactor for wastewater treatment (e.g., see Cui et al., 2003). It should be pointed out that the price of polymeric membrane dropped sharply in the last 5±10 years, which favoured the low flux, large membrane area approach. Although the concept of critical flux helped to shift the design methodology, techniques for the quantitative determination of critical flux need to be developed. Important issues such as prediction of critical flux, its dependence on solution parameters, or hydrodynamics and its transients all need to be addressed in future research. So far, most of fouling research has been performed with simulated solution, with the focus on understanding the basic mechanisms involved. Some research results have shown that a significant difference occurred when real biological solutions were used. The wider acceptance of membrane technology in the evergrowing field of biotechnology requires much more high-quality experimental work in this area. However, there is no doubt that continued efforts to develop new membranes, new modules, and optimal process design and operation will certainly consolidate the dominant roles of membranes in haemodialysis, cell harvest, mammalian cell culture and bioproducts processing. Emerging new applications in biosensors, microsampling and tissue culture have also shown great potential.

18.7 References Aimar, P., Taddei, C., Lafaille J. P. and Sanchez, V. 1988. Mass transfer limitations during ultrafiltration of cheese whey with inorganic membranes. J. Membr. Sci., 38: 203±221. Akhtar, S., Hawes, C., Dudley, L., Reed, I. and Stratford, P. 1995. Coatings reduced the fouling of microfiltration membranes. J. Membr. Sci., 107: 209±218.

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Part IV

Surface interactions and in-vivo studies

19

Bioactive 3D scaffolds in regenerative medicine: the role of interface interactions

J R J O N E S and L L H E N C H , Imperial College London, UK

19.1 Introduction Life expectancy is increasing as healthcare and technology improves, but not all body parts can maintain their function with the ageing process. Tissues such as bone and cartilage are needed to support the ageing body even though the cells that produce the tissues become less active with age. The heart, lungs, kidneys and liver have to operate for much longer than ever before. This chapter describes how biomedical materials and implant techniques are used in the clinic for tissue repair and how the interface reactions of different materials affect the repair process. The successes and shortcomings of current state of the art techniques are reviewed. The chapter then discusses how some of these materials may provide the potential for regeneration of tissues to their original state and function, rather than replacement. The surface reactions of these materials and the interactions of proteins with the next generation of biomedical materials are discussed from the results of in vivo and in vitro experiments. The use and design of scaffolds to guide tissue regeneration are reviewed and the role of the interface in the stimulation of tissue growth is discussed, focusing on bone and lung regeneration.

19.2 The need for biomedical materials and implants 19.2.1 Bone Bone is a natural composite of collagen (polymer) and bone mineral (ceramic). Collagen is a triple helix of protein chains, a complex structure that has high tensile and flexural strength and provides a framework for the bone. Bone mineral is a crystalline calcium phosphate ceramic (hydroxy apatite, HA) that provides stiffness and the high compressive strength of the bone. The two most important types of bone are cortical and cancellous bone. Cortical bone is a dense structure with high mechanical strength and is also known as compact bone. Cancellous or trabecular bone, also called spongy bone, is an internal porous supporting structure of a network of struts (trabeculae) enclosing large voids (macropores).

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The mechanism for natural bone generation/regeneration in the body is the secretion of extracellular matrix by osteogenic cells (osteoblasts), which have developed (differentiated) from stem cells. The extracellular matrix is collagen type I, which mineralises to form bone mineral, creating a composite of orientated collagen fibrils and HA. The bone is remodelled in response to its local loading environment by the body. Osteoclasts are cells that resorb old bone and bone that is not required (i.e. not under any load), while osteoblasts lay down new bone. Disease and damage can inhibit this process. Bone disease The most common bone disease is osteoporosis, which causes a loss of bone density and affects everyone as they age. The density and strength of the bone decreases because bone resorption occurs faster than new bone is produced. The struts of the trabecular bone, which is present in the ends of long bones such as the femur or within the confines of the cortical bone in short bones, are most affected by osteoporosis. The disease eventually leads to fracture of bones especially in the hip, wrist, knee and spine. At present, when osteoporotic fracture occurs in knees and hips joint replacement is often required. Critical sized defects When minor damage is done to a bone, the bone can repair itself by the activity of osteogenic cells. However, if the defect is in excess of a critical diameter or volume, the bone cannot repair itself. Such defects can result from trauma or from the removal of cancerous or diseased tissue. Graft implants (transplants) or synthetic bone filler materials are currently used to repair critical size bone defects.

19.2.2 Lung The structure of a lung is such that it has a very high surface area for gas exchange. The trachea (wind pipe) divides in to two bronchi, each of which enter one of the two lungs. The bronchi then branch into smaller bronchi and bronchioles before ending with the alveoli (air sacs), which provide a site for gaseous exchange with the arteries and veins. Primary pulmonary hypertension is a progressive and fatal disease, typified by a pulmonary arterial pressure being above 25 mm Hg at rest and at sea level (compared to 15 mm Hg for a normal adult). The disease develops in patients with congenital cardiovascular abnormalities, patients with sickle cell disease and in a group of diseases summarised under the rubric of pulmonary hypertension (Tuder, 1998). There are many lung diseases caused by the by-products of modern lifestyle, such as smoking and pollution. Emphysema and chronic obstructive pulmonary

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disease (airflow obstruction in the presence of emphysema or bronchitis) were the highest cause of death in the USA in 1998 (Snider, 2003). Emphysema is the permanent enlargement of the bronchioles and the destruction of their walls. It occurs most commonly with chronic bronchitis and impairs blood gas exchange.

19.3 Surgical procedures for bone repair There are several procedures used in the clinic for the repair of bone defects.

19.3.1 Grafts (transplants) Grafts can be taken from a donor site in the same patient (autograft), from another human donor (homograft) or from other living or non-living species (xenografts). Grafts can be used to attain the ultimate goal of restoration of a tissue to its original state and function but there are many limitations (Jones, 2001). Autografts and homografts are restricted by limited material availability and complicated multistage surgery to the detriment of the harvest site. Homografts and xenografts run the risk of disease transmission. Therefore there is a great demand for synthetic substitutes specially designed and manufactured to repair bone and cartilage (Langer, 1993).

19.3.2 Total joint replacements The most common joint replacements are the knee and hip, which are used when the fracture occurs due to a loss of bone density (osteoporosis). A total hip replacement prosthesis is comprised of a stainless steel or titanium alloy femoral stem, an alumina ball and an alumina or high density polyethylene acetabular cup as an articulating surface for the joint. An example of a total hip replacement is the Charnley total hip replacement, which is heralded as one of the most successful surgical inventions as it restores mobility to the patient. However, long-term monitoring of 20,000 Charnley joints revealed that after 25 years implantation 24% required revision surgery (Berry, 2002). Loosening of the prosthesis is a common cause of failure of the replacement joint. The interface of the material with the bone plays a large role in the survivability of the implant. The Charnley joint is made from materials that are characterised as being bio-inert. No material is completely inert on implantation, but the only response to the implantation of bio-inert materials is encapsulation of the implant by fibrous tissue (scar tissue). Therefore there is no direct bond between the femoral shaft of the prosthesis and the bone of the original femur. This means that bone cement has to be used to prevent movement of the prosthesis, which would cause damage to the host bone. However, the bone cement (polymethyl methacrylate), as well as the polyethylene articulating surface, often produces wear particles under loading.

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These particles have been found to produce inflammatory response from the body, triggering resorption of the host bone and therefore further loss of bone density. The most common reason for failure is aseptic loosening of the femoral stem, where bone resorption occurs due to a mismatch in the Young's modulus of the bone and the metal stem (stress shielding) combined with osteolysis due to the wear debris. A prosthesis is required that can bond to bone and have similar mechanical properties of the host bone and minimises creation of wear debris by eliminating interfacial micro-motion between the prosthesis and the host bone.

19.3.3 Bioactive materials in bone repair When implanted into the body, bioactive materials stimulate a biological response from the body such as bonding to tissue (Hench, 1993). There are two classes of bioactive material; class B bioactive materials bond to hard tissue (bone) and stimulate bone growth along the surface of the bioactive material (osteoconduction). Examples of class B bioactive materials are synthetic hydroxyapatite (HA), tri-calcium phosphate ceramics ( -TCP) and porous titanium oxide (titania). Class A bioactive materials not only bond to bone and are osteoconductive but they are also osteoproductive, i.e., they stimulate the growth of new bone on the material away from the bone/implant interface and can bond to soft tissues such as gingival (gum) and cartilage. Examples of class A bioactive materials are bioactive glasses. Both bioactive glasses and synthetic HA have been used clinically as bone filling agents in a particle form. -TCP is an osteoconductive material that is also resorbable in the body and is usually used in conjunction with synthetic HA to improve the resorbability of the HA in applications such as the filling of bone defects left by cysts, sinus floor augmentation and bone cements. Synthetic HA Synthetic HA has also been used to coat the metal femoral stem of hip replacement prostheses to obtain a bond to the host bone, however survivability studies are so far only medium term and have shown similar results to that of the Charnley prosthesis (Ducheyne, 1980; Bradley, 2000). Bioactive glasses A certain composition of melt-derived bioactive glass (46.1% SiO2, 24.4% Na2O, 26.9% CaO and 2.6% P2O5, in mol), called BioglassÕ is used in the clinic as a treatment for periodontal disease (PerioglasÕ) and as a bone filling material (NovaboneÕ) (Hench, 1991; Fetner, 1994). BioglassÕ implants have also been used to replace damaged middle ear bones, restoring the hearing of thousands of patients (Wilson, 1995).

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19.3.4 Interface reactions of bioactive materials in solution The mechanism of bone bonding to bioactive materials is thought to be due to the formation of a carbonate substituted hydroxyapatite layer (HCA) on the surface of the materials after immersion in body fluid. This layer is similar to the apatite layer in bone and therefore a strong bond can form. The layer forms quickest on the class A bioactive materials (Hench, 1991). Bioactive glasses After immersion in body fluid, bioactive glasses undergo a dissolution process that is instrumental in the formation of the apatite layer, which in the case of bioactive glasses is a carbonated hydroxyapatite layer (HCA). The mechanism of formation of the HCA layer on bioactive glasses is shown in Fig. 19.1. Stage 1 involves the exchange of cations between the glass and the solution. Cations such as Na+ and Ca2+ are present in the glass as network modifiers, which means they disrupt the silica network by bonding to it via non-bridging oxygen bonds. Cations such as H+ and H3O+ are present in the body fluid. Stage 2 is the break up of the silica-based glass network, i.e., the -O-Si-Obridging oxygen bonds. These two stages are dissolution processes; therefore these glasses are not only bioactive but are also resorbable in the body. The dissolution of the silica network causes the formation of silanol (Si-OH) groups along the glass surface, which act as nucleating sites for a calcium phosphate rich layer as Ca2+ and PO43- groups migrate to the glass surface both from within the glass and from the body fluid. The formation of this layer does not explain why these glasses are osteoproductive. Osteoproduction is now thought to be

19.1 Suggested reaction stages of bioactive glasses after immersion in body fluid.

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due to the dissolution products of the glasses and will be discussed in section 19.6.4. There are two types of bioactive glasses; the traditional melt-derived glasses such as BioglassÕ, and those produced by the sol-gel process. Sol-gel derived bioactive glasses tend to be more bioactive and resorb quicker than melt-derived glasses of similar compositions. This is due to sol-gel glasses having a textural porosity that is inherent to the sol-gel process, which increases the specific surface area by two orders of magnitude compared to a melt-derived glass of a similar composition (Sepulveda, 2002a). The porosity not only increases the surface area for cation exchange and network dissolution, but it also exposes many silanol groups and nucleation sites to the solution. Synthetic HA When synthetic HA is immersed in body fluid a surface layer forms. The layer consists of carbonate-apatite microcrystals of no preferred orientation with a calcium-phosphate ratio different to the bulk (LeGeros, 1993). The mechanism for the formation of this layer is thought to be due to the partial dissolution of the ceramic, releasing calcium and phosphate ions, which combine with electrolytes from the body fluid such as CO32ÿ and saturate the solution with respect to HA and precipitate on the surface of the HA forming the HCA crystals (Brown, 1999). -TCP -TCP is believed to bond to bone and then dissolve, being replaced by new bone and is classed as a resorbable ceramic. The resorption is thought to be due not only to a dissolution mechanism but also due to phagocytosis, i.e., resorption by the action of macrophage cells. Macrophages are cells that encapsulate and break down small particles of foreign material (Jarcho, 1981). When -TCP is immersed in body fluid no reactive layer forms on the surface.

19.4 Surgical procedures in lung repair The treatment for patients with parenchymal or vascular pulmonary disease is transplantation. Organ transplantation has saved the lives of many patients with organ failure and the patients often return to live active lives. However, the allografts (donor organs) are not only in short supply, with 20% of patients dying on the waiting list for a transplant (Hosenpud, 2000) but there is also a risk the patient will reject the new organs and immunosuppressant drugs have to be constantly administered. The lung has proven to be one of the most difficult organs to transplant. There are also possible complications after transplantation, such as the frequent occurrence of bronchiolitis obliterans, which leads to progressive

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dysfunction of the pulmonary allograft and eventually to respiratory failure. The continued use of immunosuppressant drugs gives rise to infection and possible malignancy as well as the side effects of the individual drugs, an example of which is enhancement of osteoporosis. The use of porcine xenografts (Cooper, 2000) and the development of artificial mechanical lungs (Hattler, 2003) are being investigated but are not yet used clinically. However some of the disadvantages of the transplant would still exist. The ability to grow a replacement lung in a laboratory from the cells of the patient is very desirable but is several decades away from clinical trials.

19.5 A new direction: regenerative medicine 19.5.1 Introduction to regenerative medicine All present-day implants lack three of the most critical characteristics of living tissues: (i) the ability to self-repair; (ii) the ability to maintain a blood supply; and (iii) the ability to modify their structure and properties in response to environmental factors such as mechanical load. All implants have a limited lifespan so it is proposed that a shift in emphasis from replacement of tissues to regeneration of tissues is required to satisfy this growing need for very long-term tissue repair. Regenerative medicine is a general term used to describe techniques that are being developed to regenerate diseased or damaged tissues to their orginal state and function using the body's own biological repair mechanisms. The techniques combine biology, materials science and biomedical engineering to achieve long-term repair and replacement of failing human tissues and organs (Berthiaume, 2000; Senuma, 2000; Sheridan 2000). Two different strategies can be employed: (i) in vitro construction of bioartificial tissues from cells harvested from the patient or from cell lines (tissue engineering), and (ii) in vivo alteration of cell growth and function (tissue regeneration) (Berthiaume, 2000). In some tissue engineering applications, scaffolds are used as templates for tissue growth (Davies, 2000; Langer, 1993). There is the potential for stem cells to be extracted from a patient, seeded on an appropriate 3D scaffold in vitro, where they will be given the biological signals to proliferate and differentiate, and then stimulated to proliferate and differentiate to form specific tissues that mimic the complex structure and physiological behaviour of natural tissues, e.g., the tissue will grow outside the body in bioreactors and develop a 3D structure, ready for implantation (Ohgushi, 1999; Takezawa, 2003). The use of the patient's own cells will eliminate problems with rejection. Tissue regeneration techniques involve the direct implantation of a scaffold (seeded with cells or not) into a defect to guide and stimulate tissue repair in situ. In both cases the scaffold should resorb (dissolve) as the tissue grows, leaving no trace of damage or surgery (Freyman, 2001). In any application of scaffolds, the material choice and design is important and in fact critical to success.

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19.5.2 An ideal scaffold for regenerative medicine To be able to regenerate trabecular bone, a construct is required that will mimic the structure of trabecular bone and stimulate new bone growth when cultured with osteogenic cells (osteoblasts). An understanding of the in vitro and in vivo behaviour of scaffolds must be obtained. An ideal scaffold is one that mimics the extracellular matrix of the host tissue so that it can act as a template in three dimensions onto which cells attach, multiply, migrate and function. The criteria for an ideal scaffold for bone regeneration are that it: · acts as template for tissue growth in three dimensions, i.e., it has an interconnected pore network with pores with diameters in excess of 100 m for cell penetration, tissue ingrowth, vascularisation and nutrient delivery to the centre of the regenerating tissue (Lu, 1999; Okii, 2001); · is made from a material that is biocompatible and bioactive, i.e., it bonds to the host tissue without the formation of scar tissue; · exhibits a surface texture that promotes cell adhesion, adsorption of biological metabolites (Hench, 1997). · influences the genes in the bone generating cells to enable efficient cell differentiation and proliferation; · resorbs at the same rate as the tissue is repaired, with degradation products that are non-toxic and that can easily be excreted by the body, for example, via the respiratory or urinary systems; · is made from a processing technique that can produce irregular shapes to match that of the defect in the bone of the patient and that can be adapted for mass production; · exhibits mechanical properties sufficient to be able to regenerate tissue in the particular application such as bone in load-bearing sites; · has the potential to be commercially producible to the required ISO (International Standards Organisation) or FDA (Food and Drug Administration) standards; · can be sterilised and maintained as a sterile product to the patient. These are the ideal criteria for a versatile scaffold, however, not all the criteria may have to be fulfilled for all applications.

19.6 Bone regeneration Osteoporosis eventually causes the loss of the trabeculae (struts) of trabecular bone (section 19.1.1). Therefore, a scaffold is required that mimics the structure of trabecular bone and can guide growth of new bone. For in situ bone regeneration, where the scaffold is implanted directly into a bone defect, mechanical properties of the scaffold are critical and the elastic modulus and strength of the porous material should match that of the natural bone. If a

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scaffold with a modulus much lower than the host bone is implanted into a loadbearing site the scaffold will fracture. If the modulus of the scaffold is much higher than bone the load will be transmitted through the scaffold instead of the bone (stress shielding), causing bone resorption rather than bone regeneration (Berry, 2002). For tissue engineering applications only the mechanical properties of the final tissue engineered construct is critical. Recently, resorbable polymeric scaffolds have been developed with porous structures similar to trabecular bone, however, polymers cannot yet bond to bone and do not stimulate genes in bone cells. Polymers have a Young's modulus much lower than bone and cannot be used directly in load-bearing applications and there is also concern that the acidic degradation products of the scaffolds will degrade cells and growth factors (Holy, 2003). HA, TCP and bioactive glasses have been used successfully in the clinic as bone filler materials in powder form (Hench, 1991; Gibson, 2002), however, the challenge is to develop them into 3D scaffolds with the properties listed above. The following sections will focus on that challenge.

19.6.1 Bioactive ceramic scaffolds The simplest way to generate porous scaffolds from ceramics such as HA, is to sinter particles, preferably spheres of equal size (Ota, 1997). As sintering temperature increases, pore diameter decreases and mechanical properties increase as the packing of the spheres increases. Mechanical properties can be increased further by hot isostatic pressing (HIPing), which will also decrease the pore diameter (Yin, 1999). Many other techniques can be used to produce porous ceramics, such as the freeze drying and the use of polymer beads or other organic agents that can be burnt out leaving pores. However, such techniques do not produce pore networks that have pores large enough or interconnected enough for tissue engineering applications. Direct foaming Highly porous ceramics can be produced through foaming of particulate suspensions or colloidal sols to obtain pores in the range of 20 m up to 1±2 mm. The incorporation of bubbles is achieved by injection of gases though the fluid medium, mechanical agitation, blowing agents, evaporation of compounds or by evolution of gas by in situ chemical reaction (Sepulveda, 1999; Ortega, 2003). A surfactant is generally used to stabilise bubbles formed in the liquid phase by reducing the surface tension of the gas-liquid interfaces (Ortega, 2001). Surfactants are macromolecules composed of two parts, one hydrophobic and one hydrophilic (Rosen, 1989). Owing to this configuration, surfactants tend to adsorb onto gas-liquid interfaces with the hydrophobic part being expelled from the solvent and a hydrophilic part remaining in contact with the liquid. This behaviour lowers the surface

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tension of the gas-liquid interfaces, making the foam films thermodynamically stable, which would otherwise collapse in the absence of surfactant. The most successful method for foaming bioactive ceramics is the gel-casting method. Gel-casting Sepulveda (1999, 2000) used the gel-casting method to produce macroporous bioactive hydroxyapatite (HA) ceramics. Suspensions of HA particles, water and dispersing agents and organic monomers (6wt% acrylate/diene), were foamed by vigorous agitation with the addition of a surfactant under a nitrogen atmosphere. In situ polymerisation of the monomers was initiated and cross-linking occurred, forming a 3D polymeric network (gel). The polymerisation process was initiated using ammonium persulphate and a catalyst (TEMED, N,N,N',N'Tetramethyl ethylene diamine) before casting. Porous samples were then sintered to provide mechanical strength and to burn out the organic solvents. Total porosity could be controlled by the surfactant concentration in the slurry, producing pores of maximum diameter of 100±200 m. Theoretically, gel-casting could be applied to melt-derived bioactive glass powders. However, such glasses undergo surface reactions on contact with solutions to produce an HCA surface layer and it is desirable to have control over the reaction before a scaffold is ready for clinical use.

19.6.2 Melt-derived bioactive glass scaffolds Producing scaffolds from melt-derived bioactive glasses such as the commercially available 45S5 BioglassÕ composition (Hench, 1991) would provide a quick route to clinical use. However, it has proven difficult to produce scaffolds with interconnected pore networks. Sintering 45S5 BioglassÕ particles mixed with camphor (C10H16O) produced scaffolds with pore diameters in the range 200±300 m (Livingston, 2002). Burning out the camphor created the large pores. However, the total porosity was just 21% as there were large distances between pores, therefore interconnectivity was low. Foaming was also investigated. Foaming of melt-derived bioactive glasses Melt-derived bioactive glass scaffolds have been produced by using a dilute H2O2 solution to foam a slurry of BioglassÕ 45S5 powder, which was then sintered. Pore diameters were in the range 100±600 m. However, the pores were few in number and were in the form of orientated channels of irregular diameter running through the glass, implying that interconnectivity was low (Yuan, 2001). Due to the limited success of producing high porosity melt-derived bioactive glasses, attention has turned to sol-gel derived bioactive glasses.

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19.6.3 Sol-gel derived bioactive glass scaffolds The sol-gel process is an alternative method to produce bioactive glasses (Hench, 1997). An advantage of gel-derived bioactive glasses over the meltderived materials is that they exhibit a mesoporous (pores with diameters in the range 2±50 nm) texture. This texture enhances bioactivity and resorbability of the glasses and creates anchorage sites for cells such as osteoblasts (Li, 1990; Sepulveda, 2002a). However, these glasses are not yet clinically available. The sol-gel process involves the production of a colloidal solution (sol) of Si-OH groups by the hydrolysis of alcoxide precursors, such as tetraethyl orthosilicate (TEOS) in excess water under acidic catalysis. Simultaneous polycondensation of Si-OH groups to form siloxane (Si-O-Si) bonds will continue after hydrolysis is complete, beginning the formation of the silicate network. As the network connectivity increases, viscosity increases and eventually a gel forms. The gel is then subjected to carefully controlled thermal processes of ageing (60 ëC) to strengthen the gel, drying (130 ëC) to remove the liquid by-product of the polycondensation reaction and thermal stabilisation/ sintering to remove organics from the surface of the gel (600±800 ëC). A chemically stable glass is then produced (Hench, 1990). Porous sol-gel derived bioactive glass scaffolds have been developed in many ways. The sintering of gel-derived glass powders (Balas, 2001), phase separation (Nakanishi, 2000) and freeze-gelation (Statham, 1998) produced pores too small (<5 m) for tissue engineering applications. Many authors have cast silica based sol-gel glasses around removable templates such as latex or polystyrene spheres (Lebeau, 2000; Khramov, 2001; Yan, 2001). Once the sol has gelled and templates are removed, voids are left in the gel. The templated pore networks were homogeneous and interconnected, due to the close packing of the spheres, but the pore size, which is at present approximately 0.5 m (Yan, 2001), is limited by the size of the organic spheres. Only foaming processes have produced sol-gel derived glass scaffolds with macropores (>10 m in diameter) that are suitable for tissue regeneration. Foaming sol-gel derived bioactive glasses Generally, the foaming of sol-gel derived glasses involves the introduction of bubbles into the sol prior to gelation. The gelation process then sets the bubbles, leaving stable pores. Two main methods have been used to introduce the bubbles into the sol. Both methods use surfactants. Foaming agents The first foaming technique uses a foaming agent that can be added to generate bubbles within a solution. Freon (CCl3F) droplets (dispersed in methanol) have

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been added to the sol (Wu, 1990). Freon has a boiling point of 23.8 ëC, therefore gas bubbles are produced above this temperature. As the mixture was heated above the boiling temperature an anionic surfactant (sodium dodecyl sulphate, SDS) was added to lower the surface tension of the sol. The viscosity of the mixture increased due to a polymerisation reaction, which caused a viscosity increase in the sol. The rate of this reaction was controlled by adjustment of pH with H2SO4 and bubble formation occurred simultaneously with an increase in viscosity and eventually gelation, which stabilised the bubbles permanently. An increase in Freon concentration and incubation temperature caused porosity and cell size interconnectivity to increase. A porous silica structure was obtained with porosities of 55±90% and corresponding mean cell diameters of 30± 1000 m. The bending strength (3 point bend) of the silica foams ranged from 2 MPa at a porosity of 86% to approximately 8 MPa at a porosity of 66%. Direct foaming Sol-gel derived bioactive glasses have been directly foamed by vigorous agitation, with the aid of surfactants, to produce scaffolds that fulfil many of the criteria of the ideal scaffold (Sepulveda, 2002b). The process has similarities to the gel-casting process. The standard sol-gel process as described above was followed, but it was modified in the following way. During the process, a mixture of anionic and non-ionic surfactants was added to the sol and the sol was foamed under vigorous agitation. The surfactants lower the surface tension of the sol and stabilise bubbles created by air entrapment. A gelling agent (hydrofluoric acid, HF) was added to the sol to increase the gelation rate to a few minutes. As gelation occurred, the bubbles were stabilised permanently. The highly porous foam was then heat treated in the same way as a normal sol-gel glass. Figure 19.2 shows an SEM micrograph of a typical foam of the 70S30C (70mol% SiO2, 30mol% CaO) composition, with macropores with diameters of up to 600 m and interconnected pore diameters (black holes) of up to 300 m. The pore structure closely mimics that of trabecular bone. The properties of the pore network can be controlled at each stage of the process, e.g.. by the type and concentration of catalyst and surfactant, the glass composition and the process temperature (Jones, 2003, 2004a). Due to the nature of the sol-gel process the scaffolds can be produced in many shapes, which are determined simply by the shape of the casting mould or cut to a required shape. The only ideal scaffold criterion not addressed is the matching of mechanical properties of the scaffolds to bone for in situ bone regeneration applications. Foams optimised by sintering have a compressive strength of ~2.5 MPa similar to that of trabecular bone (2±10 MPa) (Jones, 2004b), however, the toughness and tensile strength of the foams is much less than that of bone. The mechanical properties of these foams should be sufficient for tissue engineering applications, where bone would be grown on a scaffold in the laboratory before implantation.

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19.2 Scanning electron micrograph of a typical pore structure of a bioactive glass foam scaffold.

19.6.4 In vitro effects of bioactive glasses on bone regeneration Primary human osteoblast cells, harvested from the heads of femurs removed during total hip replacements have been seeded in vitro on to the surfaces of the foamed glasses of the 58S composition (Gough, 2004). The cells attached, spread, proliferated and secreted bone extracellular matrix, which mineralised after ten days of culture, forming bone nodules. No bone nodules were observed on tissue culture plastic after two weeks of culture (Gough, 2004). Perhaps the most significant result from this study was that mineralised bone nodules formed in the pores of the scaffold even when a basic tissue culture medium was used. For bone nodules to form on tissue culture plastic, the basic osteoblast culture medium of D-MEM supplemented with fetal bovine serum, non-essential amino acids and antibiotics must be further supplemented with ascorbic acid, dexamethasone and -glycerophosphate (Coelho, 2000). A reason as to why the mineralisation and bone nodules form without supplement could be the effect of the bioactive scaffolds on the genes of the cultured osteoblast cells. Influence of surface interactions on the genes in cells Although development of the HCA surface layer is useful, it is not the critical stage of reaction for bone regeneration (Hench, 2002). Figure 19.3 shows the suggested biological stages for bone regeneration of bioactive glasses. The sequence follows on from the events in Fig. 19.1, but concentrates on the effects of the glass on the cells within bone and bone marrow. The mechanisms of the

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19.3 Suggested mechanism for bone regeneration of the surface of reacted bioactive glass surfaces.

critical stages 3, 4 and 5 (Fig. 19.3) are now beginning to be understood. The effect of BioglassÕ versus a tissue culture plastic control on the cell cycle of primary human osteoblasts was investigated (Xynos, 2000a; Hench, 2000). A eucaryote cell cycle has four phases: G0 (the cell rests), G1 (the cell grows and performs its normal metabolism), S (DNA synthesis and chromosome duplication) and G2 (second growth phase preparing for cell division) and M (mitosis, cell division). Once the cells have divided the daughter cells restart the cycle. A critical increase in mass and synthesis and activation of various growth factors is required to pass between steps of the cell cycle and the local chemical environment must be optimised for transition from each stage to the next. If the transition does not occur, programmed cell death occurs (apoptosis). The percentages of cells in the specific phases of the osteoblast cell cycle were quantified using various cell biology methods (Xynos, 2000a,b). The results showed that the dissolution products of 45S5 bioactive glass upregulated seven families of genes in osteoblasts that regulate osteogenesis and the production of growth factors, including the genes responsible for mitosis (Xynos 2000a,b). Also, the percentage of osteoblasts in the S and G2 phases of the cell cycle, after two days of culture, were 100% greater for the cells cultured on the bioactive glass substrate compared to the tissue culture plastic. This has important implications for bone regeneration, especially for the elderly. As we age fewer progenitor cells pass the checkpoints in the cell cycle. Using bioactive glasses, the chemical environment stimulates only the viable cells to pass the checkpoints and to proliferate and lay down bone extracellular matrix and mineralise that matrix, creating new viable bone, i.e., bone regeneration. These findings may provide the reasons as to why certain compositions of bioactive glasses are osteoproductive. Only class A bioactive materials provide the chemical environment that enables osteoblasts to pass the checkpoints and produce rapid new bone formation in vivo.

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19.6.5 In vivo performance of scaffolds for bone repair Synthetic HA scaffolds This section will only discuss the performance of HA scaffolds that have been produced with similar structures to the bioactive glass foams and scaffolds for bone repair. Porous HA scaffolds produced by the gel-casting process have been implanted into the tibia of rabbits; bone partially filled the pores after eight weeks and there was no inflammatory response (Sepulveda, 2002c). Tamai (2002) implanted similar scaffolds produced by a gel-casting technique into rabbit femoral condyles. However, the scaffolds had few pores with diameters in excess of 100 m. After six weeks in vivo, evidence of new bone formation was observed throughout the pore network. Interestingly, the compressive strength of the scaffolds increased from ~10 MPa to ~30 MPa after nine weeks implantation. The increase in strength was due to tissue ingrowth that created a biocomposite with the ceramic scaffold. This increase in strength implies that it is not necessary to have a scaffold with a compressive strength equal to bone, as culturing cells on the scaffold in vitro may create a biocomposite that may increase the strength by a factor of three. This is an important finding that indicates the feasibility of use of in vitro tissue engineered constructs that solves the mechanical property mismatch inherent in the current generation of bioinert and bioactive materials. The disadvantages of HA scaffolds over bioactive glass scaffolds with similar morphology are that HA resorbs only very slowly, the dissolution products do not stimulate the genes in the osteogenic cells and HA is only osteoconductive as it generates bone at a slower rate than bioactive glasses (Oonishi, 1997). Porous HA scaffolds do not readily resorb in vitro, therefore they are likely to find applications as direct implants for bone regeneration, with or without cells seeded within the structures. It is therefore imperative that HA based scaffold implants exhibit mechanical properties similar to that of the host bone. Therefore, a biocomposite may have to be tissue engineered before implantation to provide that strength. Due to the slow rate of dissolution of synthetic HA it may take longer than the lifetime of the patient for the scaffold to be completely resorbed and the bone restored to its original state, but the function of the bone should be restored within months of implantation. Silicon has been found to be a major contributor to the mineralisation of bone and gene activation (section 19.6.4), which has led to the substitution of silicon for calcium into synthetic hydroxyapatite. In vivo results showed that bone ingrowth in silicon substituted HA granules was significantly greater than that into phase pure HA granules (Patel, 2002).

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Bioactive glass scaffolds There has been limited in vivo work on bioactive glass scaffolds, mainly due to the limited numbers of scaffolds that have been successfully produced (sections 19.6.2 and 19.6.3). The BioglassÕ 45S5 scaffolds with pore channels of diameters in the range of 100±600 m were implanted into the muscle tissue of dogs and were found to induce bone growth, containing osteoblasts and osteocytes, in soft tissue (Yuan, 2001). Bone formed directly on the solid surface and on the surface of the crystalline HCA layer that formed in the pores. Osteogenic cells were observed to aggregate near the material surface and secrete bone matrix, which then calcified to form bone. However, although the implants had a porosity of about 30% only 3% bone was formed. Sol-gel derived bioactive glass foam scaffolds of the 58S composition were implanted on rabbit crania. The foam scaffolds were compared to an empty defect, to 58S gel-derived powder and melt-derived BioglassÕ powder that is available commercially. This experiment was designed to investigate the potential for scaffolds in maxillofacial reconstruction, therefore the aim was to grow bone in a direction out and away from the host bone rather than to repair a critical sized defect within a bone. The powder was packed into a Gore-Tex sleeve to the same dimensions as the scaffold. The scaffolds stimulated new bone growth at a similar rate to that of the melt-derived bioactive glass powder available commercially and better than the sol-gel derived powder. Bone growth occurred most rapidly in the empty site in the first eight weeks after implantation, however, after eight weeks the bone was resorbed and after twelve weeks the amount of bone in the empty defect was substantially less than the foam and powder implants (Cook, 2003).

19.6.6 Summary of progress in bone regeneration Bioactive scaffolds seem to have high potential for the regeneneration of critical sized bone defects caused by trauma and tissue removal, reducing the need for transplants. Many techniques have been employed to produce porous ceramics. The technique that produces scaffolds that most closely mimic the structure of trabecular bone is the gel-casting process. Fewer techniques for producing porous bioactive glasses have been developed. Foaming sol-gel derived bioactive glasses produced scaffolds with similar macropore networks to the gel-cast ceramics. Both the gel-cast HA and sol-gel derived bioactive glass foams showed favourable results in both in vitro and in vivo tests for bone regeneration. However, bioactive glasses have controlled resorbability and release ions that can stimulate the genes in osteoblasts and control the cell cycle of osteogenic cells. It may be possible to slow or halt the progression of osteoporosis at an early stage using the bioactive glass as a drug delivery system. Administration of bioactive glass dissolution products may stimulate new bone

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growth and readdress the balance between bone growth and bone resorption, reducing the need for total joint replacements.

19.7 Tissue engineering of the lung Growing a pulmonary artery in vitro integrated with a tissue engineered lung lobe construct is one of the few alternatives to the use of heart-lung transplants for critically ill patients. Creating such a viable construct that is capable of supporting human life is a formidable challenge. Ten years ago such an enterprise would have been technically impossible. However, recent results from the fields of materials science and cell biology indicate that in vitro formation of three-dimensional functional organs may be feasible within the next decade or two. First, materials can be used to deliver proteins and growth factors needed for organogenesis and bioactive materials can be used to activate the regenerative potential of tissues (section 19.6.4). Secondly, the foaming of the bioactive sol-gel derived bioactive glasses can be used to produce hierarchical porous scaffolds that can direct growth of 3-D interconnected structures that mimic the natural architecture of tissues (section 19.6.3). The function of the bioactive composition will be to stimulate and control the rate and sequence of differentiation and proliferation of the tissues, whereas the role of the hierarchical bioactive matrix will be to direct formation of the threedimensional architecture of the engineered tissue. Thirdly, it has been shown that type II pneumocytes (lung cells) can be differentiated in vitro from mouse embryonic stems cells under suitable conditions (Ali, 2002). The alveolar surface of the lung is highly specialised for gaseous exchange. An adult alveolus is made up of two principal cell types known as type I and type II pneumocyte cells. Type I cells provide a thin barrier for gaseous exchange and play an active role in the regulation of lung liquid homeostasis. The surfaces of Type II cells are covered by type I cells except their apical pole, which bears blunt microvilli. After peripheral lung injury it has been shown that type II cells undergo a process of proliferation and differentiation to a type I phenotype (Wright, 1984). Hence type II cells not only secrete surfactant but are also progenitor cells that replace type I cells after their loss during injury. The lung has a very complex architecture and will eventually require a complex scaffold with a very well defined pore network. However, initial work on scaffolds has concentrated on whether sections of lung tissue can be grown on a material and the stimulating factors the cells need to lay down lung extracellular matrix.

19.7.1 Protein/surface interactions The interactions between cells and surfaces play a major biological role in cellular behaviour. Cellular interactions and artificial surfaces are mediated through adsorbed proteins (Mansur, 2000; Chin, 2000). If cells can recognise the

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proteins adsorbed in the surface of a biomaterial they can attach to it and start to differentiate and induce tissue formation or regeneration. However, if cells do not recognise the proteins as part of natural tissues, they will initiate a chronic inflammation that can result in encapsulation of the material or cell death (Mansur, 2000; Ratner, 1999). A common strategy in tissue engineering is to modify the biomaterial surface to selectively interact with a cell through biomolecular recognition events. Adsorbed bioactive peptides can allow cell attachment onto biomaterials, and allow three-dimensional structures modified with these peptides to preferentially induce tissue formation consistent with the cell type seeded, either on or within the device (Healy, 1999). Besides promoting cell-surface recognition, bioactive peptides can be used to control or promote many aspects of cell physiology, such as adhesion, spreading, activation, migration, proliferation, and differentiation (Drumheller, 2000). Laminin has special biomedical significance as it is a component of the lung extracellular matrix. Laminin consists of various combinations of a long chain ( chain) and two shorter chains ( and ) linked together to form an elongated cruciform shape (Murray, 1996; de Arcangelis, 1996; Tian, 1997). It is a major non-collagenous extra-cellular matrix protein, which is a major component of the basement membrane of most epithelial cell types, including the lung. It interacts with other basement membrane proteins, attaches cells to basement membranes and affects cell migration and phenotype. It has also been shown to play an important role in both foetal lung development (Schuger, 1990) and in the differentiation of isolated type II alveolar epithelial cells in vitro (Rannels, 1987; Matter, 1994). Cell interactions with laminin are mediated by a group of cell surface receptors called integrins (Virtanen, 1995, 1996), which are transmembrane glycoproteins consisting of heterodimers, which include different  and subunits. Laminin has binding sites for type IV collagen, heparin and integrins on cell surfaces. The collagen interacts with laminin, which in turns interacts with integrins or other laminin receptor proteins, thus anchoring the basal lamina to the cells. There are fourteen  and nine sub-units identified so far and they associate to form at least 23 different integrins. The 1 family has receptors for the ECM components like laminin. The  component reacts with the elements of the cytoskeleton such as actin. These integrin receptors are expressed differently in different species and at different stages and sites of airway development (Kriedberg, 1996; Pierce, 1998). Basal laminae are specialised areas of the extracellular matrix that surround epithelial cells amongst others. Glycoproteins containing an RGD sequence bind to laminin and are major cell attachment factors (Murray, 1996). Laminin is known to promote cell adhesion, proliferation and differentiation (Lekmine, 1999; Nomizu, 1995a,b). The protein/glass interactions are of greater importance in lung tissue engineering than in bone regeneration. Soft tissue bonding to bioactive glasses is

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thought to occur by tissue incorporation around HCA nodules (Hench, 1996). In order to enhance the properties of these materials or adapt them to soft tissue engineering, researchers have tried to incorporate growth factors and proteins that promote cell recognition and adhesion. The growth factors and proteins can then be released in a controlled manner, having a beneficial effect on tissue formation (Ducheyne, 1999; Nicoll, 1997; Sheridan, 2000). Polymeric systems for drug and protein delivery are numerous. However, there are presently no biomaterials that exhibit simultaneously controlled biomolecule release and bioactive behaviour (Healy, 1999). A goal for lung tissue engineering is to synthesise materials that release proteins in a controlled way while at the same time serve as a support for tissue ingrowth and direct organogenesis.

19.7.2 Modified 3D scaffolds Incorporation of bioactive peptides in three-dimensional scaffolds as an approach to tissue engineering is challenging because of the fragile nature, chemical complexity and geometric complexity of these macromolecules (Ratner, 1999; Liu, 1999). Several prerequisites must be satisfied in order to successfully graft bioactive molecules to produce materials that demonstrate cell-surface recognition and promote cell adhesion, proliferation and differentiation (Mansur, 2000; Ratner, 1999). A critical factor to the way that proteins or other bioactive peptides interact with the surfaces can alter their biological functionality. In order to achieve full functionality, peptides have to adsorb specifically. They also must maintain conformation in order to remain functional biologically. Chemical groups, such as amine and mercaptan groups are known to control the ability of surfaces to interact with proteins (Mansur, 2000; Healy, 1999; Drumheller, 2000). In addition, these chemical groups can allow proteinsurface interactions to occur such that the active domains of the protein can be oriented outwards where they can be maximally effective in triggering biospecific processes (Healy, 1999). Webb et al. (Webb, 1998) related that the adsorption of some proteins, such as bovine serum albumin, fibronectin and laminin, led to differential cell adhesion depending on whether the protein was adsorbed on different chemically functional groups. This finding indicates that proteins can have differential effects on cell behaviour depending on substrate surface chemistry. Healy (1999) cited various successful peptide attachments on inorganic and organic matrices utilising chemical functional groups and demonstrated that the cell behaviour at surfaces could be controlled in vitro by selective immobilisation of peptides.

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19.7.3 Surface-modified sol-gel derived bioactive glass scaffolds Matrices utilised for simultaneous tissue support and delivery of angiogenic factors must exhibit an appropriate pore structure, bioactive properties, peptide release kinetics, and controlled resorbtion rates. Lenza (2002) developed a novel protocol to modify the surface of bioactive glass foam scaffolds, using organosilane agents. These approaches may be ideal to incorporate and deliver angiogenic factors at controlled rates of release due to the gentle processing conditions utilised. The modified scaffolds have been used as carriers for albumin and laminin proteins, respectively (Lenza, 2003). These proteins were selected as molecular models for the eventual tailoring of scaffold chemistry to engineer growth of arterial and lung tissues. The 3-D architecture of the foams mimics the scale and interconnectivity of structures needed to grow a pulmonary artery and lung lobe in vitro. Bioactive foams modified with amine and mercaptan groups were

19.4 A graph of optical density of mouse lung epithelial cells after seven days of culture on different types of sol-gel derived bioactive foam scaffolds: 58S (58S composition; 60 mol% SiO2, 36 mol% CaO, 4 mol% P2O5), 58S(L) (58S foam coated in laminin), 58SA (58S foam modified with amine groups), 58SA(L) (58S foam modified with amine groups and coated with laminin), 58SM (58S foam modified with mercaptan groups), 58SM(L) (58S modified with mercaptan group and coated with laminin), SiO2 (100 mol% SiO2 foam).

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modified by adsorption of protein (laminin) onto their surfaces by soaking the foams in a PBS solution containing laminin (from human placenta) at 37 ëC. The efficiency of the process, in terms of protein loading, was higher than 99%, using methods outlined by Peterson (1977). Figure 19.4 shows a graph of the change in optical density of cells cultured on the different scaffold types, obtained by the WST-1 method. The number of cells proliferating on the laminin coated amine modified foams was higher than uncoated foams, laminin coated mercaptan modified foams and uncoated foams modified with amine and mercaptan groups. This was thought to be due to amine groups providing better attachment of the laminin. The patterned surfaces of the foams allow the interaction between laminin and the surface of the materials without promoting denaturation of the protein. The chemically functional groups on the surface of the foams act as ligands, to which the specific binding sites of the protein bond covalently (Lenza, 2003). The in vitro bioactivity (formation of the HCA layer) of the foam scaffolds following immersion in simulated body fluid at 37 ëC, was not affected by the surface modification process. The release of the laminin into simulated body fluid from the surface of the foam scaffolds was similar to the release profiles of silica from the scaffolds. This suggests that laminin release from the bioactive foams is driven by the dissolution rate of the material network. Therefore, the surface modification of the bioactive glasses with organosilanes and protein reduced the dissolution rate of the glass. The reduction of dissolution rate by protein adsorption has also been observed by Sepulveda et al. (2002a) and by Radin et al. (1997).

19.7.4 Gene activation and the lung All connective tissues contain critical amounts of bound silicon (Hench, 1996). One of the roles of biogenic silicon appears to be related to stabilisation of the extracellular components of the connective tissue matrix (Carlisle, 1986). This is especially the case for the arterial wall and the wall of blood vessels. Studies indicate that silicon forms complexes not only with the components of the connective tissue matrix, such as glycosaminoglycans, but also with proteins, collagen and elastin, all critical constituents of a tissue engineered pulmonary artery and lung construct. The presence of biogenic silicon is indicated to have an antiatheromatous function as well (Nadja, 1991). Figure 19.5 shows the effect of ageing on the concentration of biogenic silicon in rabbit and human aorta. There is a continuing decrease in the biologically fixed silicon with a concomitant increase of unbound, free silicon in plasma. Patients that exhibit sclerotic damage have lost biologically bound silicon from the aorta at a rate that is two to three times as fast as normal humans. This effect is attributed to the biologically cross-liked silicon inhibiting permeability of blood lipids to conjunctive tissues (Nadja, 1991). These findings

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19.5 Schematic graph showing the effect of ageing on the concentration of biogenic silicon in rabbit and human aorta.

indicate that it may be necessary to deliver critical amounts of silicon to a pulmonary artery-lung construct growing in vitro to enhance stable formation of the arterial wall structure. Bioactive glasses are an ideal delivery vehicle for controlled release of biogenic silicon because their rate of reaction is a function of the percentage of network-forming silicon in their composition (Hench, 2002). The rate of release of silicon can be controlled to match the rate of cell growth and organogenesis. The modified bioactive foams, described above, provide a means to deliver the optimal concentrations of soluble silicon as the extracellular matrix is formed in vitro, releasing proteins in a controlled and continuous way over a long period of time. Protein carriers cited in the literature, such as calcium phosphate coatings, release the majority of proteins within the first 48 hours of immersion (Nicoll, 1997). In contrast, the modified bioactive scaffolds released less than 10% protein in the first 24 hours of immersion and maintained a sustained release of proteins for a period as long as one month. This form of protein release is advantageous for applications that require release of bioactive molecules throughout a long time period for the full biological effect to be realised, such as for tissue formation or regeneration. These findings provide a technological foundation for molecular design of a tissue regenerated lung construct.

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19.7.5 Summary A matrix or scaffold can be designed that can satisfy several critical functions simultaneously: (i) an optimal composition that releases constituents to enhance growth, proliferation and differentiation of cells in vitro, (ii) a geometry that directs cellular formation of the correct 3-D architecture of the tissue construct, (iii) nanometer scale porosity that enables gases to permeate to the cells and maintain vitality, (iv) channels that can direct formation of a blood supply and innervation of the construct, (v) progressive resorption of the matrix or scaffold during growth so that the final product mimics a natural organ, and (vi) controlled release of proteins that enhance cell proliferation and differentiation. None of the presently used polymeric materials provide these complex functions in vitro. Future developments based upon molecular design and sol-gel processing of hierarchical bioactive inorganic matrices offer the potential to achieve all these functions within a single material system.

19.8 Conclusion The future of regenerative medicine lies in obtaining a full understanding of how materials can stimulate cells to carry out self-repair. The effects of different materials on the cell cycles of different human cell phenotypes should be investigated. Scaffolds should be designed from materials that have the potential to control the cell cycle, such as bioactive glasses and silicon substituted synthetic hydroxyapatite. The scaffold design also needs to be tailored at a biomolecular level specific for each application. For example, bone tissue engineering applications may require a scaffold that releases critical concentrations of osteogenically active ions to control the osteoblast stemcell cycle. On the other hand, superior mechanical properties are a priority for a scaffold designed for direct implantation in a load-bearing site. Bioactive ceramic-polymer composites or hybrid scaffolds will have to be investigated for load-bearing bone regeneration applications or for soft tissue engineering applications where flexibility of the scaffold is required. The scaffold properties should also be tailored for each patient. In order to mimic the exact morphology and structure of the scaffold required for a specific defect, X-ray computer tomography (CT) scans can be taken of the defect. The data can be input to computer aided design (CAD) files and rapid prototyping techniques will be used to synthesise scaffolds with architectures dictated by the CT scan. This technique may be useful in the creation of a scaffold suitable for the growth of a lung lobe. The technology and biomolecular science base now exists for the engineering and regeneration of most types of tissues and organs.

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19.9 Acknowledgements Royal Academy of Engineering, the EPSRC, Lloyds Tercentenary Foundation and the MRC (UK).

19.10 References Ali N (2002), Edgar A J, Samadikuchaksaraei A, Timson C M, Romanska H M, Polak J M, Bishop A E, `Derivation of type II alveolar epithelial cells from murine embryonic stem cells'. Tissue Eng, 8 (4), 541±550. Balas F (2001), Arcos D, Perez-Periente J, Vallet-Regi M, `Textural properties of SiO2CaO-P2O5 glasses prepared by the sol-gel method'. J Mate. Res, 16 (5), 1345±1348. Berry D J (2002), Harmsen W D, Cabanela M E, Morrey M F, `Twenty-five-year survivorship of two thousand consecutive primary Charnley total hip replacements', J Bone Jt Surg (Am), 84A (2), 171±177. Berthiaume F (2000), Yarmush M L, `Tissue engineering', in Bronzino J D, The biomedical engineering handbook. Boca Raton, CRC Press LLC, 1556±1566. Bradley J G (2000), Andrews C M, Lee K, Scott C A, Shaw D. `Furlong hydroxyapatite coated hip prosthesis versus the Charnley cemented hip prosthesis- A prospective randomised study'. Key Eng Mat, 192, 1013±1020. Brown P (1999), Martin R I, `An analysis of hydroxyapatite surface layer formation', J Phys Chem B, 103, 1671±1679. Carlisle E (1986), `Biological silicon'. In Evered D, Silicon Biochemistry, Ciba Found. Symp. 121, Chichester, J. Wiley and Sons, 123±136. Chin J A (2000), Slack S M, `Biomaterials: protein-surface interactions', in Bronzino J D, The biomedical engineering handbook. Boca Raton, CRC Press LLC, 1597±1608. Coelho M J (2000), Trigo Cabral A, Fernandes M H, `Human bone cell cultures in biocompatibility testing. Part I: osteoblastic differentiation of serially passaged human bone marrow cells cultured in a-MEM and in DMEM', Biomaterials, 21, 1087±1094. Cook R (2003), 58S sol-gel Bioglass: a study of osteoproductive, interfacial and handling properties using new microscopic techniques. PhD thesis. University of London. Cooper D K C (2000), Lanza R P. Xeno ± the promise of transplanting organs into humans. New York, Oxford University Press. Davies J E (2000), Bone engineering, Toronto, EM2 incorporated. de Arcangelis A (1996), Neuvil A, Boukamel R, Lefebvre O, Kedinger M and SimonAssmann P. `Inhibition of laminin a1-chain expression leads to alteration of basement membrane assembly and cell differentiation', J Cell Biol, 133, 417±430. Drumheller P D (2000), Hubbell J A, `Surface immobilisation of adhesion ligands for investigations of cell-substrate interactions', in Bronzino J D, The biomedical engineering handbook. Boca Raton, CRC Press LLC, 1583±1592. Ducheyne P (1980), Hench L L, Kagan A, Martens M, Burssens A, Mulier J C, `The effect of hydroxyapatite impregnation of skeletal bonding of porous coated implants' J Biomed Mater Res. 14, 225±237. Ducheyne P (1999), Qiu Q, `Bioactive ceramics: the effect of surface reactivity on bone formation and bone cell function'. Biomaterials, 20, 2287±2303. Fetner A E (1994), Hartigan M S, Low S B, `Periodontal repair using Perioglas in nonhuman primates: Clinical and histologic observations', Comp Cont E Dent, 15 (7),

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architecture by dual templating', J Mater Chem, 10, 2105±2108. LeGeros R Z (1993), LeGeros J P, `Dense Hydroxyapatite', in Hench L L, Wilson J, An Introduction to Bioceramics, Singapore, World Scientific, 139±180. Lekmine F (1999), Lausson S, Pidoux E, Segond N, Roos B, Treilhou-Lahille F, Jeanne N, `Influence of laminin substratum on cell proliferation and CALC 1 gene expression in medullary thyroid carcinoma C cell lines', Mol Cellr Endocrinol, 157, (1±2), 181±189. Lenza R F S (2002), Jones J R, Vasconcelos W L, Hench L L, `Surface-modified 3D scaffolds for tissue engineering'. J Mat Sci: Mat Med, 13, 837±842. Lenza R F S (2003), Jones J R, Vasconcelos W L, Hench L L, `In vitro release kinetics of proteins from bioactive foams', J Biomed Mater Res, 67A, 121±129. Li P (1990), Zhang F, `The electrochemistry of a glass-surface and its application to bioactive glass in solution', J Non-Cryst Sol, 119, 112±116. Liu D M (1999), Chen I W, `Encapsulation of protein molecules in transparent porous silica matrices via an aqueous colloidal sol-gel process', Acta Mater, 18, 4535±4544. Livingston T (2002), Ducheyne P, Garino J, `In vivo evaluation of a bioactive scaffold for bone tissue engineering', J Biomed Mater Res, 62, 1±13. Lu J X (1999), Flautre B, Anselme K, Hardoiun P, Gallur A, Deschamps M, Thierry B, `Role of interconnections in porous bioceramics on bone recolonisation in vitro and in vivo'. J Mat S-M M, 10, 111±120. Mansur H S (2000), Vasconcelos W L, Lenza R F S, OreÂfice R L, Reis E F, Lobato Z P, `Sol-gel silica based networks with controlled chemical properties'. J. Non-Cryst. Sol. 273, 109±115. Matter M L (1994), Laurie G W, `A novel laminin E8 cell adhesion site is required for lung alveolar formation in vitro', J cell Biol, 124, 1083±1090. Murray R K (1996), Granner D K, Mayes P M, Rodwell V W, Harper's biochemistry. Hartford, Connecticut, Appleton & Lange. Najda J (1991), Gminski J, Drozdz M, Flak A, `The effect of silicon (Si) on lipid parameters in blood serum and arterial wall'. Biol Trace Elem Res, 31, 235±247. Nakanishi K (2000), `Porous Gels made by phase separation: recent progress and future directions'. J Sol-Gel Sci Technol, 19, 65±70. Nicoll S B (1997), Radin S, Santos E M, Tuan R S, Ducheyne P, `In Vitro release kinetics of biologically active transforming growth factor a1 from a novel xerogel carrier', Biomaterials, 18, 853±859. Nomizu M (1995a), Weeks B S, Weston C A, Kim W H, Kleinman H K Yamada Y, `Structure-activity study of a laminin a1 chain active peptide segment Ile-Lys-ValAla-Val (IKVAV)', FEBS Lett, 365, (2±3), 217±231. Nomizu M (1995b), Kim W H, Yamamura K, Utani A, Song S Y, Otaka A, Roller P P, Kleinman H K, Yamada Y. `Identification of cell binding sites in the laminin a1 chain carboxyl-terminal globular domain by systematic screening of synthetic peptides', J Biol Chem, 270, (35), 20583±20590. Ohgushi H (1999), Caplan A I, `Stem Cell Technology and Bioceramics: From cell to Gene Engineering', J Biomed Mater Res, 48, 913±927. Okii N (2001), Nishimura S, Kurisu K, Takeshima Y, Uozumi T, `In vivo histological changes occurring in hydroxyapatite cranial reconstruction ± Case report', Neurol Med, 41 (2), 100±104. Oonishi H (1997), Kutrshitani S, Yasukawa E, Iwaki H, Hench LL, Wilson J, Tsuji E, Sugihara T, `Particulate Bioglass compared with hydroxyapatite as a bone graft

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substitute', Clin Orthop Relat Res. 334, 316±325. Ortega F S (2001), Sepulveda P, Innocentini M D M, Pandolfelli V C, `Surfactants: A necessity for producing porous ceramics', Am Ceram Soc Bull, 80 (4), 37±42. Ortega F S (2003), Valenzuela F A O, Pandolfelli V C. `Gelcasting ceramic foams with alternative gelling agents', Mater Sci F, 416-4, 512±518. Ota Y (1997), Kasuga T, Abe Y, `Preparation and compressive strength behavior of porous ceramics with b-Ca(PO3)2 fiber skeletons', J Am Ceram Soc, 80 (1), 225± 231. Patel N (2002), Best S M, Bonfield W, Gibson I R, Hing K A, Damien E, Revell P A, `A comparative study on the in vivo behavior of hydroxyapatite and silicon substituted hydroxyapatite granules', J Mat Sci Mat Med, 13 (12), 1199±1206. Pierce R A (1998), Griffin G L, Mudd M S, Moxley M A, Longmore W J, Sanes J R, Miner J H, `Senior R M Expression of laminin a3, a4, a5 chains by alveolar epithelial cells and fibroblast', Amer J Respir Cell Mol Biol, 19, 237±244. Peterson G L (1977), `A simplification of the protein assay method of Lowry et al. which is more generally applicable', Analytical Biochemistry, 83, 346±356. Radin S (1997), Ducheyne P, Rothman B, Conti A, `The effect of in vitro modelling conditions on the surface reactions of bioactive glass'. J Biomed Mater Res, 37, 363±375. Rannels S R (1987), Yarnell J A, Fisher C S, Fabisiak J P, Rannels D E, `Role of laminin in maintenence of type II pneumocyte morphology and function', Am J Physiol 253, C835±C845. Ratner B D (1999), Shi H, `Recognition templates for biomaterials with engineered bioreactivity', Curr Opin Sol St M, 4, 395±402. Rosen M J (1989), Surfactants and Interfacial Phenomena, 2nd edn, New York, John Wiley & Sons, 277±303. Schuger L (1990), O' Shea K S, Nelson B B, Varani J, `Organotypic arrangement of mouse embryonic lung cells on basement membrane extract: involvement of laminin', Development, 110, 1091±1999. Senuma Y (2000), Franceschin S, Hilborn J G, `Bioresorbable microspheres by spinning disk atomisation as injectable cell carrier from preparation to in vitro evaluation', Biomaterials, 21, 1135±1144. Sepulveda P (1999), Binner J G P, `Processing of Cellular Ceramics by Foaming and in situ Polymerisation of Organic Monomers', J Eur Ceram Soc, 2059±2066. Sepulveda P (2000), Binner J G P, Rogero S O, Higa O Z, Bressiani J C, `Production of porous hydroxyapatite by the gel-casting of foams and cytotoxic evaluation', J Biomed Mater Res, 50, 27±34. Sepulveda P (2002a), Jones J R, Hench L L, `In vitro dissolution of melt-derived 45S5 and sol-gel derived 58S bioactive glasses', J Biomed Mater Res, 61 (2), 301±311. Sepulveda P (2002b), Jones J R, Hench L L, `Bioactive sol-gel foams for tissue repair', J Biomed Mater Res, 59 (2), 340±348. Sepulveda P (2002c), Bressiani A H, Bressiani J C, Meseguer L, Konig B, In vivo evaluation of hydroxyapatite foams', J Biomed Mater Res, 62 (4), 587±592. Sheridan M H (2000), Shea L D, Peters M C, Mooney D J, `Bioabsorbable polymer scaffolds for tissue engineering capable of sustained growth factor delivery'. J Cont. Rel, 64, 91±102. Snider G L (2003), `Emphysema', in Banner N, Polak J M, Yacoub M, Lung Transplantation, Cambridge University Press, 39±53.

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Statham M J (1998), Hammett F, Harris B, Cooke, R G, Jordan, R M, Roche A J, `Netshape manufacture of low-cost ceramic shapes by freeze-gelation', J Sol-Gel Sci Technol, 12, 171±175. Takezawa T (2003), `A strategy for the development of tissue engineering scaffolds that regulate cell behaviour', Biomaterials, 24, 2267±2275. Tamai N (2002), Myoui A, Tomita T, Nakase T., Tanaka J, Ochi T, Yoshikawa H. `Novel hydroxyapatite ceramics with an interconnective porous structure exhibit superior osteoconduction in vivo', J Biomed Mater Res, 59, 110±117. Tian M (1997), Hagg T, Denisova N, Knusel B, Engvall E, Jucker M. `Laminin-a2 chainlike antigens in CNS dendritic spines', Brain Res, 764, (1±2), 28±38. Tuder R M (1998), Lee S D, Cool C D, `Histopathology of pulmonary hypertension'. Chest; 114, 1S±4S. Virtanen I A (1995), Tani T, Back N, Happola O, Laitinen L, Kiviluoto T, Salo J, Burgeson R E, Lehto V P, Kivilaakso E, `Differential expression of laminin chains and their integrin receptors in human gastric-mucosa', Am J Pathol, 147 (4), 1123± 1132. Virtanen I A (1996), Laitinen T, Tani P, Paakko L A, Laitinen R, Burgeson E, Lehto V P, `Differential expression of laminins and their integrin receptors in developing and adult human lung', Am J Respir Cell Mol Biol, 15, 184±196. Webb K (1998), Hlady V, Tresco P A, `Relative importance of surface wettability and charged functional groups on NIH 3T3 fibroblast attachment, spreading, and cytoskeletal organisation', J Biomed Mater Res, 41, 422±430. Wilson J (1995), Douek E, Rust K, `Bioglass Middle Ear Devices: 10 Year Clinical Results', in Hench L L, Wilson J, Greenspan D C, Bioceramics 8, Oxford, Pergamon, 239±245. Wright N A (1984), Alison M A, `Cell proliferation in respiratory epithelia', in Wright N A, Alison, M A, The Biology of Epithelial Cell Populations. Oxford University Press, 1068±1078. Wu M (1990), Fujiu T, Messing G L, `Synthesis of cellular inorganic materials by foaming sol-gels'. J Non-Cryst Sol, 121, 407±412. Xynos I D (2000a), Hukkanen M V J, Batten J J, Buttery L D K, Hench L L, Polak J M, `Bioglass 45S5 stimulates osteoblast turnover and enhances bone formation in vitro: implications and applications for bone tissue engineering', Calc Tiss Int, 67 (4), 321±329. Xynos I D (2000b), Edgar A J, Buttery L D K, Hench L L, Polak J M, `Ionic products of bioactive glass dissolution increase proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression and protein synthesis', Biochem. Biophys Res Comm, 276, 461±465. Yan H (2001), Zhang K, Blandford, C F, Francis L F, Stein A, `In vitro hydroxycarbonate apatite mineralisation of CaO-SiO2 sol-gel glasses with a three-dimensionally ordered macroporous structure', Chem Mater, 13, 1374±1382. Yin S (1999), Uchida S, Fujishiro Y, Ohmori M, Sato T, `Preparation of porous ceria doped tetragonal zirconia ceramics by capsule free hot isostatic pressing', Brit Ceram Trans, 98 (1), 19±23. Yuan H (2001), de Bruijn J D, Zhang X, van Blitterswijk C A, de Groot K, `Bone Induction by porous glass ceramic made from Bioglass (45S5)', J Biomed Mater Res, 58 (3), 270±276.

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Intravascular drug delivery systems and devices: interactions at biointerface K S R A O , Nebraska Medical Center, USA, A K P A N D A , National Institute of Immunology, India and V L A B H A S E T W A R , Nebraska Medical Center, USA

20.1 Introduction Drug delivery systems and biomedical devices have undergone major metamorphosis over the years, and still continue to be an ever-changing arena. As we learn more about the human body and diseases, understand the complications associated with the currently used drug delivery systems and devices, and as the discovery of new drug molecules or novel biomaterials continues, scientists are striving to modify and develop formulations and devices that are less invasive, more effective, biocompatible, and yet economical. For the purpose of drug delivery, biomaterials are used either to conjugate or encapsulate drugs to protect them from enzymatic degradation, to release them slowly for sustained drug action, or to alter biodistribution to enhance their therapeutic potential. Similarly, biomaterials are used in the fabrication of lifesaving prostheses and intravascular devices. These biomaterials invariably come in close contact either with the vasculature, cells and tissue, or body fluid. `Biointerface' means the surface where biomaterials come in contact with the biological tissue or fluid. Interactions of biomaterials at the biointerface with vascular systems such as blood or blood vessels are critical because these interactions can trigger a series of events that can be life-threatening. Therefore, investigating the biocompatibility of biomaterials at the biointerface has become an integral part of drug delivery research and device design. The present chapter focuses on the biointerface and its significance in intravascular drug delivery and device systems.

20.2 Biomaterials and biointerface A biomaterial is any substance (other than a drug) or combination of substances, synthetic or natural in origin, which can be used for any period of time, as a whole or as part of a system, to treat, augment, or replace any tissue, organ, or function of the body.1 Although various intravascular medical devices enjoy wide applications in medical technology, medical device related side effects,

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often occurring at the biointerface due to their close contact with blood, require investigation to ensure their safe use. In many instances a cascade of events is initiated with protein adsorption at the biointerface, which can then lead to thrombosis, or activation of the immune and complement systems of the body (Fig. 20.1). These reactions are primarily responsible for the rejection of these devices. Therefore, one of the stringent requirements of biomaterials for their use in vascular devices and intravascular drug delivery systems is their biocompatibility at the biointerface. The cellular response to the biointerface depends on the characteristics of biomaterials and their surrounding biological environment. In an attempt to minimize the risk associated with the use of biomaterials, the general consensus is to modify the biointerface so that the devices or systems can amalgamate with the tissue/body fluid. An alternative approach is to make them invisible to the body's defense mechanism. Blood-biomaterial interactions are more sensitive than others, and hence are considered a strong indicator of the biocompatibility of biomaterials for various biomedical applications. Modifications that are commonly carried out at the biointerface to make biomaterials more compatible fall into three categories namely, morphologic, physicochemical and biological.1 The reactions observed at the biointerface depend on the surface modification of

20.1 Schematic representation of reactions occurring at the biointerface. Diagram shows the interfacial properties that determine the success and failure of the intravascular device. The following steps occur due to (A) placement of a vascular device (biomaterial) into the vascular system, (B) various blood components like blood cells and platelets are deposited at the blood/ biomaterial interface. 1, 2, 3 represent events that occur due to the interaction of vascular devices with the system. This includes (C) thrombus formation, (D) coagulation and fibrinolysis, (E) activation of immune system. Occurrence of these events can lead to failure of the vascular device.

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the device as well as on the morphologic characteristics that include texture and the degree of roughness. These properties influence downstream cellular reactions such as protein adhesion, and interaction with blood components such as platelets. Physicochemical modifications are performed in order to alter the biological response by adjusting the hydrophilicity/hydrophobicity of polymers as well as by modifying the biointerface with functional groups or growth factor peptides to facilitate cell adhesion and cell growth.

20.3 Intravascular drug delivery systems Drug delivery systems can range from the macro scale (>1 mm) to the micro scale (100±0.1 m) or nanoscale (100±1 nm) sizes.2 These systems can alter drug distribution and kinetics, and in some cases serve as a drug reservoir in the target tissue to sustain the pharmacologic effect of the drug. Some small-scale drug delivery systems include liposomes, micelles and polymeric colloidal systems such as nanoparticles. Large-scale systems comprise drug releasing implants, vascular grafts, and stents. We have principally emphasized intravascular drug delivery systems used in cardiovascular medicine (e.g. stents), but similar issues are relevant to other intravascular devices and systems. Intravascular drug delivery could be for systemic or local drug effects, the latter especially when the disease process is a localized phenomenon. Local delivery is sometimes important for drugs that are potent but have a narrow therapeutic window. Further, the localized therapeutic dose of a drug could be significantly higher than that which can be achieved via systemic administration without toxic effects. In such cases, localized drug delivery is a better option.

20.4 Nanoparticles as an intravascular delivery system Nanoparticles are submicron drug delivery systems, generally formulated using biodegradable polymers. These can be used to achieve localized vascular drug delivery as well as to sustain a systemic drug effect. A drug delivery system in the form of nanoparticles offers the advantage of being taken up by cells and capillaries, so that the drug can accumulate at the target site. Proteins, peptides and genes can be effectively incorporated in the polymeric nanoparticles in order to prevent the clearance of these agents by the immune system of the body. Drugs or macromolecules can be adsorbed, entrapped or covalently attached to the surface of nanoparticles.3 The other main advantage of using nanoparticles as drug delivery systems is the efficiency with which their surface can be modified to achieve a therapeutic goal. In our studies, we have used biodegradable nanoparticles formulated using poly (D,L-lactide-co-glycolide) (PLGA) and poly (lactide) (PLA) polymers; these are biocompatible and FDA approved.4

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20.4.1 Nanoparticles as a reservoir for systemic drug effects The challenge in medicine is the delivery of drugs into the blood circulation and to maintain levels for a desired period of time for a better therapeutic outcome and to avoid repeated administration for better patient compliance. Nanoparticles can serve as a reservoir in the blood, and can release the encapsulated therapeutic agent slowly. In order for nanoparticles to be effective for systemic delivery, it is essential that they remain well dispersed in the blood after administration, which can otherwise lead to the formation of emboli in the pulmonary capillary bed. The second issue is to avoid their rapid clearance from the circulation by hepatic macrophages and the Kupffer cells.5 Opsonization serves to eliminate nanoparticles at a faster rate from the blood circulation. Several approaches have been investigated to modify particle surfaces to prevent the opsonization of nanoparticles and their rapid clearance by the reticuloendothelial system (RES) of the body.6 Gref et al.7 have shown that nanoparticles formulated with polyethylene glycol (PEG) modified PLA/PLGA polymers have an outer surface with PEG. These PEG modified nanoparticles demonstrated increased blood circulation time for the encapsulated drug. In their studies, it was found that 66% of unmodified particles were eliminated by the liver in a period of only five minutes following their intravenous administration in mice, whereas even after two hours, only 30% of 20 kD PEG-coated nanoparticles were sequestered by the liver. Due to the PEG coating on the surface, opsonization of the nanoparticles is prevented. Several other groups since then have started adopting a similar strategy to modify the nanoparticle surface for intravascular as well as for extravascular drug delivery.8 On similar lines, polysialylation of colloidal drug carrier systems has been proposed to enhance the systemic circulation time of therapeutic agents. Polysialic acid is a polysaccharide, found in the body attached to certain cell-adhesion molecules, where it has an anti-adhesive function. There are no known natural receptors for polysialic acid, which thus acts as nature's stealth molecule. Polysialylation of proteins and peptides has been shown to prolong their systemic circulation time, and the above effect has been attributed to a reduced renal clearance and uptake by the RES.9 A further example of surface modification that has been used to attain prolonged systemic circulation times is the use of amphiphilic block copolymers (ABCs). An important example of ABC polymers includes Pluronic block copolymers (PBCs). PBCs contain varying lengths of ethylene oxide (EO) and propylene oxide (PO) blocks. These copolymers are increasingly used in drug delivery systems due to their unique architecture and low toxicity. Therapeutic drugs can be incorporated in these systems within a core that is formed by PO, and the EO side blocks that protrude into the aqueous solution impart increased systemic circulation time.10 At or above the threshold of critical micelle concentration (CMC), these ABCs have the ability to form

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polymeric micelles. Depending upon the length of polyethylene oxide that is used during the formulation, systemic circulation time as well as the uptake by the RES can be modified as desired.11 Nanoparticles with surface adsorbed PBCs also have prolonged systemic circulation times.12 These `long circulating particles' thus act as a reservoir from which drug can be released into the vascular compartment in a regular and controlled manner. This strategy is useful when a greater systemic drug exposure and reduced clearance are desirable. Thus, one can potentially enhance the half-life of a drug using long-circulating nanoparticles. Avgoustakis et al.13 have demonstrated that the intravenous administration of PLGA-mPEG nanoparticles of cisplatin in BALB/c mice resulted in prolonged cisplatin residence in the systemic blood circulation. Another type of polymer, known as dendrimers, are nanometer sized highly branched polymers that have a tree-like appearance.14 The surface of dendrimers can be modified and also drugs can be incorporated in their central cavity. Due to their nanometer size, dendrimers are being explored as an intravascular drug delivery system.

20.4.2 Nanoparticles for vascular tissue delivery Nanoparticles can be used for localized delivery of a drug to a particular segment of an artery to achieve a sustained drug effect. Our interest is in investigating nanoparticles for localized drug and gene delivery to inhibit restenosis following angioplasty.15 We have previously shown localized delivery of nanoparticles in rat carotid and porcine coronary models of restenosis. The localization efficiency of nanoparticles in arterial tissue was dependent upon various factors such as the surface properties of nanoparticles,16 particle size, conditions of infusion (concentration and volume)17 and catheter design.18 Arterial uptake was greater for smaller sized nanoparticles and increased with an increase in nanoparticle concentration in the infusion suspension. We have also demonstrated sustained drug retention in the infused section of an artery and inhibition of restenosis in a rat carotid artery model with a single-dose localized administration of dexamethasone-loaded nanoparticles. 19 So in a clinical scenario, an interventional cardiologist, immediately following a standard angioplasty procedure, can infuse a suspension of drug-loaded nanoparticles in the target segment of the artery using a standard cardiac infusion catheter. Localized delivery of nanoparticles could provide sustained therapeutic drug concentration in the target artery that could inhibit the hyperplasia, and since the total dose administered will be low because of localized drug delivery, the therapy could be without the risk of systemic toxicity. Other investigators are taking advantage of an increased permeability of the injured artery following angioplasty for targeting drug-loaded micellar nanosystems via systemic administration.20 Increased permeability is due to severe inflammation and increased expression of mitogenic agents (VEGF,

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bradykinins, etc.) in the injured artery. Uwatoku et al.21 studied the effect of NK91, a dexamethasone-conjugated nanoparticle carrier system, which when administered intravenously, inhibited vascular smooth muscle cell proliferation and reduced the formation of neointima in the rat carotid artery model of restenosis. In addition to local vascular drug delivery, endothelium could be the target for drug therapy. Dysfunctional endothelium has been implicated in several inflammatory and degenerative disorders including cardiac, respiratory, and arthritic conditions.22,23 In order to improve the status of diseased endothelium, different strategies have been used to target drugs to such endothelium using, for example, drug conjugates and immunoliposomes. In our studies, it has also been shown that nanoparticles are effectively taken up by endothelial cells, without compromising cell viability.24 Further, we have demonstrated that nanoparticles can act as an intracellular depot, releasing the drug slowly and intracellularly, a most critical feature in chronic disease conditions.25 Therefore, nanoparticles can be a useful drug carrier system for targeting therapeutic agents to the endothelium. Within the endothelium, ICAM-1 (intercellular adhesion molecule-1), which is implicated in inflammatory conditions, is a good therapeutic target for drug delivery. Muro et al.26 have shown that catalase containing anti-ICAM conjugated nanoparticles imparted antioxidative action to the endothelial cells. In an in vivo study, it was demonstrated that vascular targeting of catalase, reduced oxidative stress in rat pulmonary endothelium. 27 This property of nanoparticles could be beneficial in the case of oxidative stress where ICAM-1 is found to be upregulated. The surface property of nanoparticles could play an important role in their intracellular delivery. `Stealth' nanoparticles, which are formulated to remain in the systemic circulation, may have reduced uptake by endothelial cells. The above issue of the surface property of nanoparticles and their influence on cellular uptake has been studied in our recent studies with PLGA/PLA nanoparticles. Polyvinyl alcohol (PVA) is routinely used as an emulsion stabilizer during the formulation of nanoparticles. While formulating nanoparticles, the hydrophobic portion of PVA anchors itself into the matrix of nanoparticles. A fraction of PVA remains associated with the surface of nanoparticles despite several washings, thereby forming an interface. This residual amount of PVA has been shown to affect the interfacial properties of nanoparticles, including their hydrophilicity/hydrophobicity, which also had an effect on their cellular uptake. It was observed that nanoparticles formulated using 5% PVA as an emulsifier has a greater amount of surface associated PVA than those formulated with 2% PVA. It was shown that the cellular uptake of nanoparticles in smooth muscle cells was significantly lower in the case of nanoparticles formulated with 5% PVA as compared with the uptake of those formulated using 2% PVA (Fig. 20.2a).28 These studies hold interest, because it can be argued that along with the desired goal of achieving a longer circulation time, it is also important that the

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20.2 (a) Effect of residual PVA on the nanoparticle (NP) uptake in VSMCs. Nanoparticles were incubated with VSMCs at 100 g/well in the presence or absence of serum, data as mean  SEM (n ˆ 6). SM, serum containing medium; SFM, serum free medium*, # P < 0:05 compared to uptake of corresponding 2% PVA nanoparticle group. Reprinted with permission from ref. 28. (b) Effect of PVA concentration on transfection of nanoparticles in breast cancer (MCF-7) cells. Cells (35,000 per well in 24-well plate) were incubated with nanoparticles (444 g/ml/well) for 1 day, and then the medium in wells was replaced with fresh medium (without nanoparticles). Medium was changed on every alternate day thereafter, and luciferase protein levels were determined at one, three, five, and seven days post-transfection in MCF-7 cell line. Identical protocol was used to determine transfection of plasmid DNA or plasmid DNA + PVA. Data shown as mean  SEM, n ˆ 6. Figure legend represents concentration of PVA used as an emulsifier. Reprinted with permission from ref. 29.

delivery system releases the drug intracellularly. The latter issue can be exemplified from our recent studies in which gene expression for nanoparticles with a higher amount of surface associated PVA was significantly lower than that for nanoparticles with lower surface associated PVA, despite similar DNA loading and DNA release kinetics (Fig. 20.2b).29 However, the above issue could be addressed if long-circulating nanoparticles are modified with a targeting ligand specific for endothelial cells. These ligands have been shown to selectively target genes to the endothelium. Particular examples in this direction include gene-containing nanoparticles that are coupled to  3targeting ligand.30 These ligand specific nanoparticles have been shown to selectively target genes to the angiogenic blood vessels of a tumor. In the same study, it was shown that nanoparticles conjugated to a mutated raf gene prevented angiogenesis in response to growth factors.30 These studies, thus demonstrate that vascular endothelium can be the target in various disease conditions.

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20.5 Stents Stents represent an important intravascular device system, commonly used in patients that have undergone interventional procedures such as balloon angioplasty. A stent is a small lattice shaped metal tube that is deployed inside the lumen of an injured artery segment following angioplasty. It acts as a scaffold that prevents abrupt closure as well as the elastic recoiling of the injured artery. However, stenting also leads to in-stent restenosis (i.e. reobstruction of the artery), which is the result of multiple factors including injury to the vessel wall, endothelial dysfunction and the inflammatory process due to the stent material itself. Vascular smooth muscle cells migrate and proliferate inside the lumen of a stented artery with time that causes reobstruction. Mechanical stress on the artery due to a deployed stent and the biocompatibility of biomaterials used in the fabrication of stents also play an important role in the process of instent restenosis.31 As stents are in direct contact with the vasculature, their compatibility at the biointerface with the vessel wall as well as with the blood is essential. Thrombosis of the stented artery is a major issue, and hence patients need to be on antithrombotic therapy, at least during the first six months following stent deployment. Thus, research is focused on fabricating stents that are more biocompatible with vascular tissue. Stents coated with a polymer loaded with an antithrombotic agent (e.g., heparin), although able to prevent thrombosis, do not have any significant effect on inhibition of restenosis, suggesting that a dual drug therapy may be required, one to prevent thrombosis and the other to prevent hyperplasia.32,33 Metal stents coated with an inert polymer material have demonstrated reduced platelet deposition following stent placement. The polymer coating probably reduces electronegativity and the associated corrosion of the metal. Therefore, platelet activation and also the resulting aggregation is prevented.34 The above observation thus suggests that the biointerface is an important factor in the outcome of stents in clinical studies. Since systemic drug therapies for the treatment of in-stent restenosis have been largely disappointing, drug-eluting stents have come into existence. Stents are coated with a polymer containing a therapeutic agent that has antiproliferative activity. The eluted drug is expected to prevent the proliferation of vascular smooth muscle cells, and hence inhibit restenosis. Drug-eluting stents can deposit the released drug into the target artery, thus providing targeted drug therapy.35 Although a myriad of drugs have been studied for the treatment of restenosis, the most common ones that have gained importance for therapy are, Sirolimus (Rapamycin) and paclitaxel (Taxol). These drugs have been studied extensively in clinical trails. Sirolimus eluting stents were found to be safe and efficacious two years after their implantation in human subjects,36 thus demonstrating their long-term safety. Another antiproliferative drug, paclitaxel also had an effect in reducing the incidence of restenosis.37

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Despite promising results with drug-eluting stents as compared with bare stents, these are not without adverse effects, and cannot be considered absolutely safe because they can cause intra-intimal hemorrhage and rebound hyperplasia once the coated drug is completely eluted.37 Therefore, evaluation of drugcoated stents for long-term outcomes (beyond five years) is critical. Further, restenosis rate is significantly higher in stented smaller blood vessels (< 3.0 mm) as compared to larger vessels (> 3 mm). Also, stents are not effective if the lesion is longer than the standard stent length.

20.6 Vascular grafts and catheters Vascular grafts are prosthetic tubes that serve the purpose of replacing damaged blood vessels. Synthetic vascular grafts are most commonly made from polyethylene terephthlate, also known as Dacron.38 Just like any other vascular devices, interaction of the graft surface with the vasculature can result in inflammation or formation of thrombus.39 Researchers have employed different strategies to prevent graft failure. One approach is coating of the graft with a lining of endothelial cells to improve hemocompatibility.40 Vascular grafts are commonly used in aneurysms, which is an abnormal enlargement or bulging of an artery caused by damage to, or weakness in, a blood vessel wall.41 Although aneurysms can occur in any of the body's blood vessels, they almost always form in an artery, which upon rupture can cause life-threatening bleeding. Catheters are commonly used to administer medication or nutrients in the bloodstream, sometimes extending over ten days, and also in peritoneal and hemodialysis. These catheters are associated with a substantial risk of bloodstream infection or peritonitis.42 Microorganisms usually adhere to the intraluminal or extraluminal surface of the catheter before infection of the bloodstream. Apart from taking preventive measures such as preparing skin using antiseptic solution prior to catheterization or rinsing of a catheter with a solution of antibiotic (usually an antibiotic solution is instilled within the catheter and left there for 6±12 h before removal), new strategies have been proposed such as impregnating the catheter polymer with antibiotics or modifying the catheter surface so that bacteria do not adhere to the catheter surface. Silver and silicone coated dialysis catheters have been studied as a way of reducing bacterial colonization and the occurrence of infection.43,44

26.7 Future trends To address complex problems and innovations, an interdisciplinary approach is generally required. Intravascular drug delivery systems and devices are no exception. To fabricate effective intravascular drug delivery systems or devices, an integrated approach involving polymer scientists, biologists, pharmaceutical researchers, clinicians, and biomedical engineering is necessary. Over the last

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decade, there has been better understanding of biomaterials, biointerface, and the biology of interactions. Slow but steady progress, made over the years in the area of stents, from design to drug delivery, is considered as an important development. As a result, stents are more acceptable than before, and are used in more than 50% of the patients that undergo balloon angioplasty. The next generation of stents will probably be fabricated from biodegradable polymers, which will serve the same purpose as regular stents, i.e., providing support to the injured artery, but would completely resorb with time once the injured artery has been repaired. Such stents are expected to be free of long-term complications. Similar progress is expected in intravascular drug delivery systems, especially in the area of nanotechnology for cardiovascular applications that may range from the development of cell-specific delivery systems to biosensors for the early detection of pathologies. Also, tissue regenerating smart biomaterials could be a reality in the near future. In all these applications, the biocompatibility of biomaterials at the biointerface will remain an important determinant.

20.8 References 1. Dee K C, Puleo D A, and Bizios R. An Introduction to Tissue-Biomaterial Interactions. John Wiley and Sons, 2002. 2. LaVan D A, McGuire T, and Langer R. 2003. `Small-scale systems for in vivo drug delivery'. Nat Biotechnol, 21, 1184±91. 3. Taylor S, Qu L, Kitaygorodskiy A, Teske J, Latour R A, and Sun Y P. 2004. `Synthesis and characterization of peptide-functionalized polymeric nanoparticles'. Biomacromolecules, 5, 245±8. 4. Panyam J, and Labhasetwar V. 2003. `Biodegradable nanoparticles for drug and gene delivery to cells and tissue'. Adv Drug Deliv Rev, 55, 329±47. 5. Moghimi S M, Hunter A C, and Murray J C. 2001. `Long-circulating and targetspecific nanoparticles: theory to practice'. Pharmacol Rev, 53, 283±318. 6. Peracchia M T, Fattal E, Desmaele D, Besnard M, Noel J P, Gomis J M, Appel M, d'Angelo J, and Couvreur P. 1999. `Stealth PEGylated polycyanoacrylate nanoparticles for intravenous administration and splenic targeting'. J Control Release, 60, 121±8. 7. Gref R, Minamitake Y, Peracchia M T, Trubetskoy V, Torchilin V, and Langer R. 1994. `Biodegradable long-circulating polymeric nanospheres'. Science, 263, 1600±3. 8. Moghimi S M, and Szebeni J. 2003. `Stealth liposomes and long circulating nanoparticles: critical issues in pharmacokinetics, opsonization and protein-binding properties'. Prog Lipid Res, 42, 463±78. 9. Jain S, Hreczuk-Hirst D, Laing P, and Gregoriadis G. 2004. `Polysialylation: the natural way to improve the stability and pharmacokinetic of protein and peptide drugs'. Drug Delivery Systems & Science, 4, 3±9. 10. Oh K T, Bronich T K, and Kabanov A V. 2004. `Micellar formulations for drug delivery based on mixtures of hydrophobic and hydrophilic Pluronic block copolymers'. J Control Release, 94, 411±22. 11. Adams M L, Lavasanifar A, and Kwon G S. 2003. `Amphiphilic block copolymers for drug delivery'. J Pharm Sci, 92, 1343±55.

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12. Goppert T M, and Muller R H. 2003. `Plasma protein adsorption of Tween 80- and poloxamer 188-stabilized solid lipid nanoparticles'. J Drug Target, 11, 225±31. 13. Avgoustakis K, Beletsi A, Panagi Z, Klepetsanis P, Karydas A G, and Ithakissios D S. 2002. `PLGA-mPEG nanoparticles of cisplatin: in vitro nanoparticle degradation, in vitro drug release and in vivo drug residence in blood properties'. J Control Release, 79, 123±35. 14. Patri A K, Majoros I J, and Baker J R. 2002. `Dendritic polymer macromolecular carriers for drug delivery'. Curr Opin Chem Biol, 6, 466±71. 15. Labhasetwar V, Song C, and Levy R. 1997. `Nanoparticles drug delivery system in restenosis'. Adv Drug Del Rev, 24, 63±85. 16. Labhasetwar V, Song C, Humphrey W, Shebuski R, and Levy R J. 1998. `Arterial uptake of biodegradable nanoparticles: effect of surface modifications'. J Pharm Sci, 87, 1229±34. 17. Song C, Labhasetwar V, Cui X, Underwood T, and Levy R J. 1998. `Arterial uptake of biodegradable nanoparticles for intravascular local drug delivery: results with an acute dog model'. J Control Release, 54, 201±11. 18. Panyam J, Lof J, O'Leary E, and Labhasetwar V. 2002. `Efficiency of dispatch and infiltrator cardiac infusion catheters in arterial localization of nanoparticles in a porcine coronary model of restenosis'. J Drug Target, 10, 515±23. 19. Guzman L A, Labhasetwar V, Song C, Jang Y, Lincoff A M, Levy R, and Topol E J. 1996. `Local intraluminal infusion of biodegradable polymeric nanoparticles. A novel approach for prolonged drug delivery after balloon angioplasty'. Circulation, 94, 1441±8. 20. Hou D, Rogers P I, Toleikis P M, Hunter W, and March K L. 2000. `Intrapericardial paclitaxel delivery inhibits neointimal proliferation and promotes arterial enlargement after porcine coronary overstretch'. Circulation, 102, 1575±81. 21. Uwatoku T, Shimokawa H, Abe K, Matsumoto Y, Hattori T, Oi K, Matsuda T, Kataoka K, and Takeshita A. 2003. `Application of nanoparticle technology for the prevention of restenosis after balloon injury in rats'. Circ Res, 92, 62±9. 22. Urbich C, and Dimmeler S. 2004. `CD40 and vascular inflammation'. Can J Cardiol, 20, 681±3. 23. Kelm M. 2003. `The L-arginine-nitric oxide pathway in hypertension'. Curr Hypertens Rep, 5, 80±6. 24. Davda J, and Labhasetwar V. 2002. `Characterization of nanoparticle uptake by endothelial cells'. Int J Pharm, 233, 51±9. 25. Panyam J and Labhasetwar V. 2004. `Sustained cytoplasmic delivery of drugs with intracellular receptors using biodegradable nanoparticles'. Mol Pharm, 1, 77±84. 26. Muro S, Cui X, Gajewski C, Murciano J C, Muzykantov V R, and Koval M. 2003. `Slow intracellular trafficking of catalase nanoparticles targeted to ICAM-1 protects endothelial cells from oxidative stress'. Am J Physiol Cell Physiol, 285, C1339±47. 27. Kozower B D, Christofidou-Solomidou M, Sweitzer T D, Muro S, Buerk D G, Solomides C C, Albelda S M, Patterson G A, and Muzykantov V R. 2003. `Immunotargeting of catalase to the pulmonary endothelium alleviates oxidative stress and reduces acute lung transplantation injury'. Nat Biotechnol, 21, 392±8. 28. Sahoo S K, Panyam J, Prabha S, and Labhasetwar V. 2002. `Residual polyvinyl alcohol associated with poly (D,L-lactide-co-glycolide) nanoparticles affects their physical properties and cellular uptake'. J Control Release, 82, 105±14. 29. Prabha S, and Labhasetwar V. 2004. `Critical determinants in PLGA/PLA

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nanoparticle-mediated gene expression'. Pharm Res, 21, 354±64. 30. Hood J D, Bednarski M, Frausto R, Guccione S, Reisfeld R A, Xiang R, and Cheresh D A. 2002. `Tumor regression by targeted gene delivery to the neovasculature'. Science, 296, 2404±7. 31. Kavanagh C A, Rochev Y A, Gallagher W M, Dawson K A, and Keenan A K. 2004. `Local drug delivery in restenosis injury: thermoresponsive co-polymers as potential drug delivery systems'. Pharmacol Ther, 102, 1±15. 32. Hardhammar P A, van Beusekom H M, Emanuelsson H U, Hofma S H, Albertsson P A, Verdouw P D, Boersma E, Serruys P W, and van der Giessen W J. 1996. `Reduction in thrombotic events with heparin-coated Palmaz-Schatz stents in normal porcine coronary arteries'. Circulation, 93, 423±30. 33. De Scheerder I, Wang K, Wilczek K, Meuleman D, Van Amsterdam R, Vogel G, Piessens J, and Van de Werf F. 1997. `Experimental study of thrombogenicity and foreign body reaction induced by heparin-coated coronary stents'. Circulation, 95, 1549±53. 34. Rogers C, and Edelman E R. 1995. `Endovascular stent design dictates experimental restenosis and thrombosis'. Circulation, 91, 2995±3001. 35. Bhatia V, Bhatia R, and Dhindsa M. 2004. `Drug-eluting stents: new era and new concerns'. Postgrad Med J, 80, 13±8. 36. Sousa J E, Costa M A, Sousa A G, Abizaid A C, Seixas A C, Abizaid A S, Feres F, Mattos L A, Falotico R, Jaeger J, Popma J J, and Serruys P W. 2003. `Two-year angiographic and intravascular ultrasound follow-up after implantation of sirolimuseluting stents in human coronary arteries'. Circulation, 107, 381±3. 37. Farb A, Heller P F, Shroff S, Cheng L, Kolodgie F D, Carter A J, Scott D S, Froehlich J, and Virmani R. 2001. `Pathological analysis of local delivery of paclitaxel via a polymer-coated stent'. Circulation, 104, 473±9. 38. Chuter T A. 2002. `Stent-graft design: the good, the bad and the ugly'. Cardiovasc Surg, 10, 7±13. 39. L'Heureux N, Paquet S, Labbe R, Germain L, and Auger F A. 1998. `A completely biological tissue-engineered human blood vessel'. Faseb J, 12, 47±56. 40. Rademacher A, Paulitschke M, Meyer R, and Hetzer R. 2001. `Endothelialization of PTFE vascular grafts under flow induces significant cell changes'. Int J Artif Organs, 24, 235±42. 41. Hall S W. 2003. `Endovascular repair of abdominal aortic aneurysms'. Aorn J, 77, 631±42; quiz 45±8. 42. Crnich C J, and Maki D G. 2002. `The promise of novel technology for the prevention of intravascular device-related bloodstream infection. II. Long-term devices'. Clin Infect Dis, 34, 1362±8. 43. Trerotola S O, Johnson M S, Shah H, Kraus M A, McKusky M A, Ambrosius W T, Harris V J, and Snidow J J. 1998. `Tunneled hemodialysis catheters: use of a silvercoated catheter for prevention of infection ± a randomized study'. Radiology, 207, 491±6. 44. Schierholz J M, Rump A F, Pulverer G, and Beuth J. 1998. `Anti-infective catheters: novel strategies to prevent nosocomial infections in oncology'. Anticancer Res, 18, 3629±38.

21

Surface degradation and microenvironmental outcomes C C C H U , Cornell University, USA

21.1 Introduction The interest in biodegradable polymeric biomaterials for biomedical use has increased significantly during the past two decades due to two major advantages that non-biodegradable biomaterials do not have. First, they don't elicit a permanent chronic foreign-body reaction because they are gradually absorbed within the human body and do not permanently retain a trace of residual material in the implantation site. Second, some of them have been found recently to be able to regenerate tissues, achieving so-called tissue engineering, through the interaction of their biodegradation with immunologic cells like macrophages. Hence, surgical implants made from biodegradable biomaterials could be used as a temporary scaffold for tissue regeneration. This approach toward the reconstruction of injured, diseased or aged tissues is one of the most promising fields in the current century. The term biodegradation is loosely associated with materials that could be broken down naturally, either through hydrolytic mechanisms without enzymes or/and through enzymatic catalytic mechanisms. Other terms like absorbable, erodible, resorbable have also been used in the literature to indicate biodegradation. Although the earliest and most commercially significant biodegradable polymeric biomaterials originated from linear aliphatic polyesters like polyglycolide, polylactide and their copolymers from poly(-hydroxyacetic acids), the recent introduction of several new synthetic and natural biodegradable polymeric biomaterials extends the domain beyond this family of simple polyesters. These new commercially significant biodegradable polymeric biomaterials include poly(orthoesters), polyanhydrides, polysaccharides, poly(ester-amides), tyrosine-based polyarylates or polyiminocarbonates or polycarbonates, poly(D,L-lactide-urethane), poly( -hydroxybutyrate), poly(caprolactone), poly(amino acids), pseudo-poly(amino acids) and copolymers derived from amino acids and non-amino acids poly[bis(carboxylatophenoxy) phosphazene], and star-shaped multiarm poly--caprolactone and its hydrogels.

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The earliest, most successful and frequent biomedical application of biodegradable polymeric biomaterials has been in wound closure (Chu et al., 1997). All biodegradable wound closure biomaterials are based upon glycolide, lactide and their copolymer family. For example, polyglycolide (Dexon from American Cyanamid), poly(glycolide-L-lactide) random copolymer with a 90 to 10 molar ratio (Vicryl and Vicryl Plus from Ethicon), poly(ester-ether) (PDS and PDSII from Ethicon), poly(glycolide-trimethylene carbonate) random block copolymer (Maxon from American Cyanamid, now US Surgical), and poly(glycolide--caprolactone) copolymer (Monocryl from Ethicon). This class of biodegradable linear aliphatic polyester biomaterials is also the one most studied with regard to chemical, physical, mechanical, morphological and biological properties and in their change with degradation time and extrinsic factors such as pH, stress, superoxide etc. Some of the above materials like Vicryl have been commercially used as surgical meshes for hernia and bodywall repair. Recently, polyglycolide was electrospun into near submicron size fibers for tissue engineering (Boland et al., 2001). The next largest biomedical application of biodegradable polymeric biomaterials that is commercially satisfactory is for drug control/release devices. Some well-known examples for this application are polyanhydrides and poly(ortho-ester). Biodegradable polymeric biomaterials, particularly totally resorbable composites have also been experimentally used in the field of orthopedics, mainly as components for internal bone fracture fixation like PDS pins. However, their wide acceptance as other orthopedic implants may be limited due to their inherent mechanical properties and their biodegradation rate. Besides the commercial uses described above, biodegradable polymeric biomaterials have been experimented on as vascular grafts, vascular stents, vascular couplers for vessel anastomosis, nerve growth conduits, for augmentation of bone defects, as ligament/tendon prostheses, as intramedullary plugs during total hip replacement, anastomosis rings for intestinal surgery and stents in ureteroureterostomies for accurate suture placement. In this chapter, the emphasis will be on those commercially most significant and successful biomedical biodegradable polymers (i.e. linear aliphatic polyesters), particularly those in fiber form for wound closure, and the effects of some extrinsic microenvironments on the in vitro and in vivo degradation of the commercially most significant biodegradable polymers with the emphasis on cell and biomaterial surface interaction during the course of biodegradation. The details of the applications of this family and of other biodegradable polymeric biomaterials and their chemical, physical, mechanical, biological and biodegradation properties can be found in other recent reviews (Barrows, 1986; Kimura, 1993; Shalaby, 1994; Chu et al., 1997; Chu, 2000; Hollinger, 1995; Vert et al., 1992).

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21.2 Chemistry of synthetic biodegradable biomaterials There are basically four different types of synthetic biodegradable biomaterials that are commercially important. They are polyglycolic acid (PGA), polylactiic acid (PLA), poly-p-dioxanone (PDS), and poly--caprolactone. Most commercial biodegradable products like wound closure devices, orthopedic pins, dental regenerative membranes, and surgical meshes are made from one or a combination of some of the above four types of synthetic biodegradable biomaterials. Among these four types of synthetic biodegradable biomaterials, PGA is the commercially most successful one. Polyglycolic acid is the first and most wellknown synthetic biodegradable biomaterial, introduced commercially in the early 1970s. PGA has the simplest chemical structure in the family and that on which other synthetic biodegradable biomaterials have been built (Frazza et al., 1971; Schmitt et al., 1967; Katz et al., 1970). For wound closure purposes, PGA in suture form as DexonÕ from Davis and Geck (now part of US Surgical under Tyco International) has been one of the most familiar names in the field. There are a variety of Dexon-based wound closure biomaterials: Dexon `S', an uncoated PGA, Dexon Plus, coated with a copolymer of poly(oxyethyleneoxypropylene) and Dexon II, coated by polycaprolate to facilitate handling properties, knot performance and smooth passage through tissue. Recently, PGA-based sutures made by companies other than Davis/Geck have been introduced, such as MedifitÕ from the Japan Medical Supply Co., Safil from B. Braun Melsungen AG and PolySyn FA from Surgical Specialities. PGA is polymerized from the cyclic dimer of -hydroxyacetic acid, more commonly called glycolic acid. On heating, glycolic acids form cyclic dimers, called glycolides, more readily than other hydroxy acids due to the formation of stable, ring-strain free six-member ring anhydrides. PGA can be polymerized either directly or indirectly from glycolic acid. The direct polycondensation produces a polymer of Mn less than 10,000 because of the requirement of a very high degree of dehydration (99.28% up) and the absence of monofunctional impurities. For PGA of molecular weight higher than 10,000 it is necessary to proceed through the melt ring-opening polymerization of the cyclic dimers of glycolic acid. PGA was found to exhibit an orthorhombic unit cell with dimensions a ˆ 5.22 AÊ, b ˆ 6.19 AÊ, and c (fiber axis) ˆ 7.02 AÊ. The planar zigzag-chain molecules form a sheet structure parallel to the ac plane, and do not have the polyethylene type arrangement. The tight molecular packing and the close approach of the ester groups might stabilize the crystal lattice and contribute to the high melting point of PGA (224±227 ëC). The heat of fusion of 100% crystallized PGA is reported to be 12 KJ/mole or 45.7 cal/gram (Brandrup et al., 1975). A recent study of injection molded PGA discs reveals their IR spectroscopic characteristics (Chu et al., 1995) with four bands at 850, 753, 713, and 560 cmÿ1 that are associated with the amorphous regions of the injection-molded PGA

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discs and which could be used to assess the extent of hydrolysis. Peaks associated with the crystalline phase include those at 972, 901, 806, 627, and 590 cmÿ1. Two broad, intense peaks at 1142 and 1077 cmÿ1 can be assigned to C-O stretching modes in the ester and oxymethylene groups, respectively. These two peaks appear to be associated mainly with ester and oxymethylene groups originating in the amorphous domains. Chu et al. reported that hydrolysis could cause both of these C-O stretching modes to decrease substantially in intensity (Chu et al., 1995). The glycolide-L-lactide random copolymer suture material (Vicryl), sometimes called polyglactin 910 is also copolymerized in the same fashion as PGA. For biomedical use, the glycolide-L-lactide copolymers must have a high concentration of glycolide monomer for achieving proper mechanical and degradation properties, e.g., Vicryl sutures contain a 90/10 molar ratio of glycolic to L-lactic acid and are coated with 2±10% of a 50:50 mixture of an amorphous polyglactin 370 (a 65/35 mole ratio of lactide-glycolide copolymer) and calcium stearate. If DL- instead of L-lactide is used as the co-monomer, however, the U-shape relationship between the level of crystallinity and glycolide composition disappears. This is because polylactide from 100% DLlactide composition is totally amorphous. A relatively new block copolymer of glycolide and carbonates, such as trimethylene carbonate, has been commercialized. MaxonÕ is made from a block copolymer of glycolide and 1,3-dioxan-2-one (trimethylene carbonate or GTMC) and consists of 32.5% by weight (or 36 mole %) of trimethylene carbonate (Katz et al., 1985; Casey et al., 1984). Maxon is a poly(ester-carbonate). The polymerization process of Maxon is divided into two stages. The first stage is the formation of a middle block which is a random copolymer of glycolide and 1,3dioxan-2-one. Diethylene glycol is used as an initiator and stannous chloride dihydrate (SnCl2
2H2O) serves as the catalyst. The polymerization is conducted at about 180 ëC. The weight ratio of glycolide to trimethylene carbonate in the middle block is 15:85. After the synthesis of the middle block, the temperature of the reactive bath is raised to about 220 ëC to prevent the crystallization of the copolymer and additional glycolide monomers as the end blocks are added into the reaction bath to form the final triblock copolymer. A commercial product (ResolutÕ from W.L. Gore and Associate) consisting of a copolymer of glycolide, trimethylene carbonate and lactide has been used in dental restoration of bony defects. Resolut membrane was reported to facilitate guided bony tissue regeneration due to periodontal disease. This membrane is essentially absorbed in human body in 6±7 months, but remains substantially intact during the first 8± 10 weeks (Tonetti et al., 1998; Sanz et al., 1997). Poly-p-dioxanone (PDS) is derived from the glycolide family providing better flexibility. It is polymerized from ether-containing lactones, 1,4-dioxane2,5-dione (i.e. p-dioxanone) monomers with a hydroxylic initiator and tin catalyst (Shalaby, 1994). The resulting polymer is semi-crystalline with Tm

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about 106±115 ëC and Tg ÿ10 to 0 ëC. The improved flexibility of PDS relative to PGA as evidenced in its lower Tg is due to the incorporation of an ether segment in the repeating unit which reduces the density of ester linkages for intermolecular hydrogen bonds. Because of the less dense ester linkages in PDS when comparing with PGA or glycolide-L-lactide copolymers, PDS is expected, and has been shown, to degrade at a slower rate in vitro and in vivo. PDS having an inherent viscosity 2.0 dL/g in hexafluoroisopropanol is adequate for making monofilament sutures. Recently, an advanced version of PDS, PDSII, was introduced. PDSII was achieved by subjecting the melt-spun fibers to a high temperature (128 ëC) for a short period of time (Broyer, 1994). This additional treatment partially melts the outermost surface layer of PDS fibers and leads to a distinctive skin-core morphology. The heat employed also results in larger crystallites in the core of the fiber than the untreated PDS fiber. The tensile strength loss profile of PDSII is better than PDS. The latest glycolide-based copolymer that has become commercially successful is the Monocryl suture. It is a segmented block copolymer consisting of both soft and hard segments. The purpose of having soft segments is to provide good handling property like pliability, while the hard segments are used to provide adequate strength. The generic copolymerization process between glycolic acid and -caprolactone was reported in Japan (Fukuzaki et al., 1989, 1991). The resulting copolymers were low molecular weight biodegradable copolymers for potential drug delivery purposes. The weight average molecular weight ranged from 4,510 to 16,500 and the glass transition temperature ranged from 18 to ÿ43 ëC, depending on the copolymer composition and molecular weight. Monocryl is made from a two-stage polymerization process (Bezwada et al., 1995). In the first stage, soft segments of prepolymer of glycolide and -caprolactone are made. This soft segmented prepolymer is further polymerized with glycolides to provide hard segments of polyglycolide. Monocryl has a composition of 75% glycolide and 25% -caprolactone and should have a higher molecular weight than those glycolide/-caprolactone copolymers reported by Fukuzaki et al. for proper mechanical properties. The most unusual aspect of the Monocryl monofilament suture is its pliability as claimed by Ethicon (Bezwada et al., 1995). The force required to bend a 2/0 suture is only about 2.8  104 lb-in2 for Monocryl, while the same size PDSII and Maxon monofilament sutures require about 3.9 and 11.6  104 lb-in2 force, respectively. This inherent pliability of Monocryl is due to the presence of soft segments and its low Tg (between 15 and ÿ36 ëC).

21.3 In vitro degradation of synthetic biodegradable biomaterials The reported biodegradation studies of a variety of biodegradable polymeric biomaterials have mainly focused on their tissue biocompatibility, the rate of drug release, and loss of strength and mass. Recently, degradation mechanisms and the

590

Surfaces and interfaces for biomaterials

Table 21.1 Structural factors to control the polymer degradability. (Kimura. Y., 1993, `Biodegradable Polymers', in Biomedical Applications of Polymeric Materials, ed. T. Tsuruta, T. Hayashi, K. Kataoka, K. Ishihara and Y. Kimura, pp. 164±190. CRC Press, Inc., Boca Raton, Fla.) Factors

Methods of control

Chemical structure of main chain and side groups Aggregation state Crystalline state Hydrophilic/hydrophobic balance

Selection of chemical bonds and functional groups Processing, copolymerization Polymer blend Copolymerization, introduction of functional groups Micropores Fiber, film, composite

Surface area Shape and morphology

effects of intrinsic and extrinsic factors, such as pH (Chu, 1981a, 1982), enzymes (Chu et al., 1983; Williams, 1979; Williams et al., 1977, 1984), -irradiation (Chu et al., 1982, 1983; Campbell et al., 1981; Williams et al., 1984; Zhong et al., 1993), electrolytes (Pratt et al., 1993), cell medium (Chu et al., 1992), annealing treatment (Chu et al., 1988), plasma surface treatment (Loh et al., 1992), external stress (Chu, 1985a; Miller et al., 1984), superoxide treatment (Lee et al., 1999, 2000), and polymer morphology (Chu et al., 1989) and on the chemical means to examine the degradation of PGA fibers (Chu et al., 1985) have been systemically examined, and the subject has been reviewed (Chu, 1985b, 1991, 1995a; Chu, 2000; Hollinger, 1995). Table 21.1 is an illustration of structural factors of polymers that could control their degradation (Kimura, 1993). In this chapter, those extrinsic factors that are closely related to the biological environment (e.g., enzymes, superoxide, free radicals) and processing parameters (e.g., -irradiation) will be reviewed so that the information can be correlated with section 21.4.

21.3.1 Role of enzymes Enzymatic-aided degradation of either natural or synthetic polymers can proceed in a random or terminal mode. For example, -amylase randomly hydrolyzes C1-C4 linkages of the amylose chain and hence -maltose and glucose are the end products (Schnabel, 1981). -amylases hydrolyze the same C1-C4 linkage but at the nonreducing end of the chain so that maltose molecules are removed successively. Not all proteases hydrolyze proteins in the same way. Trypsin degrades the carbonyl groups of arginine and lysine only, while pepsin and chymotrypsin randomly attack the peptide linkages. In the case of synthetic polymers, enzymes are believed to attack at the chain ends of polymers and will be discussed later. Regardless of their purpose, all enzymes are substrate specific, and their catalytic activities relate to their chain conformation. Any

Surface degradation and microenvironmental outcomes

591

change in the chain conformation due to pH or temperature change of the medium will result in a loss of catalytic function. It has often been argued that in vivo synthetic polymer degradation may be aided by enzymes. This is based on the histological observation of the presence of macrophages immediately adjacent to degradable polymers, the principal producers of a series of lysosomal enzymes including hydrolases. It was reported that high molecular weight substrates may be degraded upon lysosomal enzymatic action in three ways (Kopecek et al., 1983): 1. 2. 3.

via endocytosis after the substrate has been captured into the cell; release of the lysosomal enzymes into both the extracellular and intracellular spaces during inflammation; by the process of autophagy.

Because the size of surgical implants is always significantly greater than biological cells, and because most polymers are insoluble in body fluids, the two broad processes of endocytosis, (i) the ingestion of particulate matters (phagocytosis) and (ii) the engulfing of small droplets of extracellular fluid (pinocytosis), and the process of autophagy appear to be inapplicable to the initial stage of biodegradation of biodegradable surgical devices and materials. Thus, the release of lysosomal enzymes into the extracellular space is the most probable process in response to the presence of biodegradable surgical implants. Such a process is especially common with older macrophage populations (chronic inflammation) (Niemi et al., 1968) and in the surrounding sites of toxic substances (Salthouse, 1976). It is important to recognize that such enzymatic activities of macrophages are mainly responsible for the degradation and absorption of natural absorbable biomaterials, such as catgut and reconstituted collagen. Whether or not such release of lysosomal enzymes from macrophages aid the degradation and absorption of synthetic biodegradable wound closure biomaterials is inconclusive. It has been mentioned previously that degradation of biopolymers with the assistance of enzymes proceeds either in a random or terminal scission mode. Schnabel postulated that, contrary to biopolymers, synthetic polymers, except polyvinyl alcohol and poly--caprolactone, are usually attacked only at the chain ends and thus degradation is affected by the molecular weight of the polymers (Schnabel, 1981). The higher the molecular weight, the smaller the number of chain ends available for enzymatic attack. These fewer chain ends in high molecular weight polymers may not be readily accessible to enzymes because they are often lost in the mass of entangled polymer matrix. Consequently, the rate of degradation should be lower as the molecular weight of polymers increases. Potts demonstrated such a molecular weight effect on the microbial degradation of hydrocarbons (Potts, 1978). A recent reported study of the effect of irradiation on the in vitro enzymatic degradation of PGA absorbable sutures by Chu et al. also suggested that the

592

Surfaces and interfaces for biomaterials

hypothesis of preferential terminal attack of absorbable polymers by enzymes may be the route for the eventual enzymatic aided absorption of absorbable biomaterials in vivo (Chu et al., 1983). In addition to the effect of the molecular weight of substrates described above, the morphology of absorbable biomaterials (i.e. crystallinity and orientation) must also bear a relationship to the enzymatic degradation of synthetic absorbable biomaterials, because crystalline domains are not accessible by enzymes which are significantly larger than water molecules. Because of their large molecular size, enzymes are too large to be able to diffuse into the interior of a synthetic biodegradable polymers. Therefore, enzymes must limit their activity to the surface or a thin surface layer of the implant unless microcracks initially existed or resulted from other types of degradation (Zaikov, 1985). Because of the relatively high hydrophilicity of absorbable biomaterials and their capability to produce lower molecular weight degradation products upon hydrolysis which could provoke additional tissue reactions (i.e., more enzymes would be released by inflammatory cells), synthetic absorbable biomaterials are expected to exhibit unique behavior toward enzyme-catalyzed biodegradations. The histochemical procedures used to study the mechanism of suture absorption (Salthouse et al., 1976) permits researchers to identify several important enzymes presented in tissues immediately adjacent to absorbable biomaterials like Vicryl absorbable sutures. These enzymes are (i) alkaline phosphatase generally associated with neutrophils and confined to the seven-day suture implant only; (ii) acid phosphatase aminopeptidase and -glucuronidase associated with macrophage and giant cells at suture sites and an exhibited higher activity between 28 and 35 days postimplantation with -glucuronidase exhibiting a lower level of intensity; (iii) nonspecific esterase frequently associated with giant cells and an activity lower than other of hydrolases like acid phosphatase, -glucuronidase and aminopeptidase; (iv) adenosine triphosphatase present from 28 to 42 days postimplantation and in areas close to suture fibers; (v) succinic dehydrogenase and lactic, isocitric and malic dehydrogenases with their highest activity in Vicryl suture between 28 and 42 days and in cells adjacent to suture fibers undergoing hydrolysis; (vi) cytochrome oxidase with its activity more conspicuous in cells bordering Vicryl fibers with a minimal activity at a distance remote from the suture site. The experimental evidence from the Salthouse study demonstrates that the primary breakdown of Vicryl is independent of cellular or enzyme activity and that the only requirement for these absorbable biomaterials to degrade is an aqueous environment. Oxidoreductase enzymes, however, appear to be associated with the metabolism of the degradation products of these biomaterials. The high activity associated with oxidative enzymes was not found in cells removed from suture implantation sites. Salthouse et al. (1976) suggested that the increase in the oxidative enzymes associated with the citric acid cycle indicates that glycolic and lactic acid degradation products are being actively

Surface degradation and microenvironmental outcomes

593

metabolized by those cells in contact with the suture filaments. Salthouse et al. concluded their enzymatic study by suggested the following in vivo scheme: Vicryl suture + Water # Glycolic and lactic acids # Macrophage oxidoreductase enzymes # Carbon dioxide and water In contrast, other investigators demonstrated that certain enzymes, under some conditions, are able to influence the degradation of PGA and polylactic acids (Williams et al., 1977; Williams, 1979). Williams et al. found that certain enzymes, under some in vitro conditions, were able to influence the degradation of 2/0 PGA sutures as shown in Table 21.2. Acid phosphatases, papain, pepsin, peptidase, pronase, proteinase-K, and trypsin had no apparent effect on PGA sutures. Ficin, carboxypeptidase-A, chymotrypsin, and clostridiopeptidase-A all produced significantly greater amounts of degradation, often increasing the rate of hydrolysis by a factor of two. Bromelain, esterase, and leucine aminopeptidase-treated PGA sutures lost all of their tensile strength after three weeks, while the untreated ones lost only 13.3% of their original value within the same period. The enzymes that did influence hydrolysis were mainly (although not exclusively) of the type (esterases) that might be expected to attack an aliphatic polyester on the basis of its molecular structure. As mentioned earlier, alteration of the structure of polymer chains by irradiation may change the susceptibility of the polymer toward enzymatic degradation. Thus, those enzymes like trypsin which show no effect on PGA initially may influence its degradation after the alteration of its physical and chemical structure by -irradiation. This possibility was examined by Chu and Williams in the in vitro study of the effect of -irradiation on the role of enzymes in PGA suture hydrolysis (Chu et al., 1983). Of the three enzymes studied (esterase, -chymotrypsin, trypsin), esterase showed the greatest enzymatic effect on the degradation of the unirradiated and irradiated PGA sutures. Irradiation made the suture more susceptible to esterase attack than was the case with an unirradiated sample, as evidenced by the lower percentage of retention of tensile breaking strength in enzymatic solutions. Trypsin's effect on PGA sutures was not observed until 20 Mrads of radiation were applied. Trypsin treated with 20 Mrad irradiated PGA sutures lost 13% more than their corresponding buffer control group after three days of immersion. Other researchers also concluded that trypsin does not accelerate the hydrolytic degradation of PGA sutures in vitro after comparing the percent loss of tensile strength of PGA sutures exposed to nonactivated canine pancreatic and enterokinase activated

2 weeks 3 weeks

5 days 7 days

Leucine aminopeptidase

5.07  0.07 4.82  0.08

4.21  0.15 3.99  0.10

4.44  0.07 4.18  0.04

3.51  0.11 1.59  0.13

3.96  0.09 3.98  0.07

(b) In solutions containing ammonium sulphate: Bromelain 2 weeks 4.40  0.07 3 weeks 4.97  0.14

Esterase

2.29  0.06

4.70  0.01 4.76  0.02 4.70 0.01

Ficin

3 weeks 4ÃÙÄ weeks 3 weeks

Clostridiopeptidase A

2.36  0.13 0.33  0.12

3.58  0.05

4.78  0.05

2 weeks

Buffer treated breaking strength (kg)  S.D.

-Chymotrypsin

Control breaking strength (kg)  S.D. 3.26  0.12

Time

(a) In preparation without ammonium sulphate: Carboxypeptidase A 2 weeks 4.45  0.06

Enzyme

12.4 13.3

15.2 60.2

10.9 19.9

51.3

49.8 93.1

25.1

26.7

% Loss

0.50  0.15 0.00  0.00

1.75  0.20 0.00  0.00

2.34  0.10 1.31  0.13

1.60  0.10

0.98  0.08 0.07  0.03

2.97  0.11

2.57  0.04

Enzyme treated breaking strength (kg)  S.D.

90.1 100.0

58.4 100.0

46.9 73.6

66.0

79.1 98.5

41.6

42.2

% Loss

Table 21.2 Effects of various enzymes on the breaking tensile strength of polyglycolic acid (Williams, D.F. and E. Mort, `Enzyme-accelerated hydrolysis of polyglycolic acid', J. Bioeng., 1(3): 231, 1977)

Surface degradation and microenvironmental outcomes

595

pancreatic juice (Mizuma et al., 1977). Chu et al.'s findings for trypsin demonstrated the hypothesis that synthetic high molecular weight polymers, which are initially resistant to enzymatic degradation, can become prone to enzymatic attack after alteration of their physical and chemical structures. Holbrook used urine as a model to examine the enzymatic effect on PGA degradation (Holbrook, 1982). He found that amylase, hyaluronidase, and urokinase failed to show enzymatic catalytic degradation, but the addition of porcine esterase into a control urine did accelerate the breakdown of PGA sutures. The suture in the esterase-treated urine had only 15% of the tensile strength (0.09 kg) of the urine control (0.61 kg) after three days. This esterase accelerated degradation of PGA was further confirmed by competitive inhibition with another ester (glyceryl triacetate-triacetin). Thus, Holbrook concluded that the destruction and fracturing of PGA sutures is more than the result of pure hydrolysis. By using a transfer between in vitro and in vivo conditions, Williams (1979) showed that there is something specific about the immediate environment in bodily implantation that influences the hydrolytic degradation of PGA. As shown in Fig. 21.1, there was a rapid initial loss of tensile strength of PGA sutures (about 10% during the first two days of postimplantation) which was not observed under in vitro conditions. Those PGA sutures which were transferred from an in vitro to an in vivo environment at day six exhibited this same rapid loss of strength immediately after the transfer.

21.1 Degradation profiles of polyglycolide sutures that were transferred between in vitro and in vivo conditions (Williams, 1979 with permission).

596

Surfaces and interfaces for biomaterials

In an attempt to identify which enzymes were responsible, those released by polymorphonuclear leukocytes during the acute inflammatory response or those released by the macrophages of the chronic response, Williams used a novel approach involving repeated implantation, removal, and reimplantation of PGA in rats for a short period of time (Williams, 1982). It was found that the reimplanted PGA, which experienced a series of acute responses, had slightly higher degradation rates than the control group, which only experienced an acute response once followed by chronic response (55.8% strength remaining in repeated implantation, removal, and reimplantation compared to 59.9% in control group at p < 0:005). The difference, however, was hardly conclusive enough to prove the hypothesis that enzymes released from cells in response to the trauma of implantation are responsible for the initial rapid loss of strength during the first few days after implantation. The possible role of enzymes in the hydrolytic degradation of synthetic absorbable biomaterials was further demonstrated by the effect of enzymatic wound cleaning surgical practice on 3/0 PGA sutures (Persson et al., 1986). Varidase, an enzymatic solution consisting of streptokinase and streptodornase, has been used as a wound cleanser in infected wounds to remove necrotic tissue, fibrin, pus, and coagulated blood (Poulsen et al., 1983). Because the cleaning procedure involves the soaking of the wound with Varidase solution, the sutures in the wounds were in close contact with the enzymatic solution. Persson et al. found that a statistically significant difference (p < 0:05) in the tensile properties of the PGA sutures was observed between saline and Varidase incubated 3/0 PGA sutures. For example, the percentage of retention of original tensile breaking strength, elongation at break, and toughness at day 15 was respectively 59%, 52%, and 29% for the Varidase incubated PGA group, while saline incubated PGA sutures retained 71%, 65%, and 47%, respectively. Similar differences in the change of mechanical properties of 3/0 PGA sutures were observed in artificially created abdominal wounds in male Wistar rats. All the reported in vivo biodegradation properties of synthetic absorbable polymers in the literature include the involvement of direct contact inflammatory cells to polymers. As described earlier, it was suggested that the close adhesion of macrophages to absorbable polymer surfaces must bear a direct relationship to the observed in vivo biodegradation property of that polymer (Salthouse, 1976). In order to isolate the effect of these biological cells from the general in vivo environment, Zhong et al. have recently reported the use of Vicryl suture as the model in a study of their biodegradation property in a controlled acellular in vivo environment (Zhong et al., 1993). Vicryl was confined within a poly(methyl methacrylate) chamber which had a pore size < 0:45 m and the chamber implanted subcutaneously in Lister rats. Because of the pore size, no biological cells were expected to penetrate into the chambers; however, biochemicals like enzymes could still diffuse into the chamber.

Surface degradation and microenvironmental outcomes

597

The SEM observations indicated that Vicryl under this controlled in vivo environment degraded faster than in vitro in PBS medium. A comparison of the surface morphological structure change of Vicryl between the in vivo chamber environment (Fig. 21.2) and in vitro PBS medium (Fig. 21.3) indicated that the in vivo chamber environment induced a faster degradation as evident in the earlier appearance of multiple surface cracks. Far more severe fragmentation

21.2 Scann ing elect ro n m icro gr aph s of 2/ 0 Vicryl s ut ures after poly(methylmethacrylate) chamber implantation subcutaneously in rats for 14 days (Zhong et al., 1993 with permission).

598

Surfaces and interfaces for biomaterials

was observed in the chamber environment at 14 days, while the same suture at a PBS environment showed sparse fragments and fewer surface cracks. This study, however, did not determine whether these two micro-environments would alter the tensile strength loss profiles.

21.3 Scanning electron micrographs of 2/0 Vicryl sutures after in vitro degradation in PBS medium for 14 days. (a)  150; (b)  480; (c)  1,000 (Zhong et al., 1993 with permission).

Surface degradation and microenvironmental outcomes

599

21.3.2 Role of free radical and superoxide ions It had been demonstrated that the biodegradation of synthetic absorbable biomaterials is closely related to macrophage activity through the close adhesion of macrophages onto the surface of absorbable biomaterials (Matlaga et al., 1980). It is also known that inflammatory cells, particularly leukocytes and macrophages are able to produce highly reactive oxygen species like superoxide (
O2ÿ) and hydrogen peroxide during inflammatory reactions toward foreign materials (Badwey et al., 1980; Devereux et al., 1991). These highly reactive oxygen species participate in the biochemical reaction, frequently referred to as a respiratory burst which is characterized by one electron reduction of O2 into superoxide via either NADPH or NADH oxidase. The resulting superoxide radicals (
O2ÿ) are then neutralized to H2O2 via the cytoplasmic enzyme superoxide dismutase (SOD). Very recently, Lee et al. and Williams et al. suggested that these reactive oxygen species may be harmful to the polymeric implant surface through their production of highly reactive, potent and harmful hydroxyl radicals
OH in the presence of metals like iron. (Lee et al., 1999, 2000; Ali et al., 1993; Zhong et al., 1994). Williams et al.'s study of absorbable biomaterials like Vicryl in the presence of aqueous free radical solutions prepared from H2O2 and ferrous sulfate raised the possibility of the role of free radicals in the biodegradation of synthetic absorbable biomaterials (Williams et al., 1991; Zhong et al., 1994). As shown below, both
OH radicals and OHÿ are formed in the process of oxidation of Fe‡2 by H2O2 and could exert some influence on the subsequent hydrolytic degradation of absorbable biomaterials. Fe‡2 + H2O2 ÿ! Fe‡3 +
OH + OHÿ

21.1

SEM morphology indicated that Vicryl in the presence of free radical solutions exhibited many irregular surface cracks at both seven and fourteen days in vitro, while the two controls (H2O2 or FeSO4 solutions) did not have these surface cracks. Surprisingly, the presence of surface cracks of Vicryl sutures treated in the free radical solutions did not accelerate the tensile breaking strength loss as would be expected. Thermal properties of Vicryl under free radical and 3% H2O2 media showed the classical well-known maximum pattern of the change of the level of crystallinity with hydrolysis time. The level of crystallinity of Vicryl peaked at seven days in both media (free radical and 3% H2O2). The time for the crystallinity peak appearance in these two media was considerably earlier than with Vicryl in conventional physiological buffer media. Based on Chu's suggestion of using the time for the appearance of the crystallinity peak as an indicator of degradation rate (Chu, 1981b), it appears that these two media accelerated the degradation of Vicryl when compared with regular physiological buffer solution. Based on their findings, Zhong et al. proposed possible routes for the role of
OH radicals in the hydrolytic degradation of Vicryl (Zhong et al.,

600

Surfaces and interfaces for biomaterials

1994). Unfortunately, the possible role of OHÿ, one of the byproducts from Fenton reagents (H2O2/FeSO4), was not considered in the interpretation of their findings. OHÿ species could be more potent than
OH towards the hydrolytic degradation of synthetic absorbable biomaterials. This is because hydroxyl anions are the sole species which attack the carbonyl carbon of the ester linkages of absorbable aliphatic polyester-based biomaterials during alkaline hydrolysis. Since an equal amount of
OH and OHÿ is generated in Fenton reagents, the observed changes in morphological, mechanical and thermal properties could be partially attributed to OHÿ ions and to
OH radicals. Besides hydroxyl radicals, the production of superoxide ions and singlet oxygen during phagocytosis has been well documented (Babior et al., 1973); but their role in the hydrolytic degradation of synthetic biodegradable polyesterbased biomaterials has remained largely unknown. Such an understanding of the superoxide ion role during the biodegradation of foreign materials has become increasingly desirable because of the advanced understanding of how the human immune system reacts to foreign materials and increased use of synthetic biomaterials for human body repair. Lee et al. recently examined the reactivity of the superoxide ion towards biodegradable biomaterials having an aliphatic polyester structure under different reaction conditions, such as temperature, time, and superoxide ion concentration (Lee et al., 1999, 2000). Due to the extreme reactivity of the superoxide ion, they observed that the effect of superoxide ion induced hydrolytic degradation of poly(D,L) lactide (PDLLA) and poly-L-lactide (PLLA) was significant in terms of their change in molecular weight and thermal properties. Superoxide ion induced fragmentation of PDLLA would result in a mixture of various species with different chain length. The significant reduction in molecular weight of PDLLA by superoxide ion was also evident in the change of thermal properties like Tg. The mechanism of simple hydrolysis of ester by superoxide ion proposed by Forrester et al. (1984) was subsequently modified by Lee et al. (2000) to interpret the data obtained from the synthetic biodegradable polymers. The exact mechanism, however, is not fully known yet; Lee et al. (2000) suggested the possibility of simultaneous occurrence of several main-chain scissions by three different nucleophilic species. In addition to PDLLA and PLLA, superoxide ions also have a significant adverse effect on the hydrolytic degradation of synthetic absorbable sutures (Lee et al., 2000). Thus there is a significant reduction in molecular weight, mechanical and thermal properties of such sutures over a wide range of superoxide ion concentrations, particularly during the first few hours of contact with these ions. For example, a PGA suture lost almost all of its mass at the end of 24 h contact with superoxide ion at 25 ëC, while the same suture would take at least 50 days in vitro buffer for complete mass loss. The surface morphology of these sutures was also altered drastically. The appearance of moon cratershaped impressions of various sizes (about 10±100 m diameter) on MonocrylÕ

Surface degradation and microenvironmental outcomes

601

21.4 Surface morphology of 2/0 Monocryl suture upon superoxide ioninduced hydrolytic degradation at 25 ëC for 24 hours. The superoxide ion concentration was 0.0025 M (Lee et al., 2000 with permission).

suture at a superoxide ion concentration > 0.005 molar (Fig. 21.4) is unique because such circular impressions were never observed during the hydrolytic degradation of all available absorbable sutures in conventional saline buffer medium or in vivo. The formation of moon-crater shaped impressions on Monocryl and Maxon sutures deviates from the conventional understanding of the anisotropic characteristic of fibers. It appears that these circular-shaped impressions started randomly on the suture fiber surface and propagated concentrically (i.e., uniformly at all angles), irrespective of the fact that all fibers are highly anisotropic. In the reported morphological study of all existing absorbable sutures in conventional buffer media (Chu et al., 1997), the most common surface morphological characteristic upon hydrolytic degradation of suture fibers is the formation of circumferential or/and longitudinal surface cracks that are consistent with the anisotropic characteristic of fibers. It is not fully understood at this stage how superoxide ion induced degradation could lead to such unusual moon-crater shaped surface morphology on Monocryl and Maxon sutures.

21.4 In vivo biodegradation of synthetic biodegradable biomaterials and cell/biomaterial surface interaction When implanted in living tissues, biodegradable biomaterials inevitably elicit a wide range of tissue reactions and cellular responses. Although inflammation due to surgical trauma subsides within seven to fourteen days post-operation, inflammatory tissue reactions due to the presence of biomaterials would persist as long as they remain present within the tissue. An excessive inflammatory reaction is undesirable because it lowers body defense mechanisms against infection, delays the subsequent proliferate phase of wound healing, and leads to inferior wound strength due to excessive scar tissue formation.

602

Surfaces and interfaces for biomaterials

There are two basic approaches to evaluating these tissue reactions; histology and enzyme histochemistry. The former evaluates cellular responses which can be assessed according to the reaction degree and area by microscopic observation of the type and population size of the responding cells, as well as edema and necrosis. The latter is particularly useful in the evaluation of in vivo absorption mechanisms of absorbable biomaterials. All histologic evaluation has some subjective component and, therefore, some type of numerical scoring system should be used in order to achieve a more independent and objective evaluation. Among the evaluation criteria, the use of Sewell's method or its modification for assessing the tissue response to implanted biomaterials has been well accepted (Sewell et al., 1955). In the Sewell method, scores for the width of the reaction zone, cellular density, cell type, subsequent edema, and necrosis are combined as illustrated in Table 21.3 (Sewell et al., 1955; Chu, 1985b). The degree of reaction is judged on the total score: 0, no reaction; 1 to 8, minimal; 9 to 24, slight; 25 to 40, moderate; 41 and above, marked reaction. The appropriateness of applying this scoring system to modern biomaterials has been questioned (Smit et al., 1991). Smit et al.'s major criticism was that the Table 21.3 Sewell score system for semi-quantitative histological evaluation of suture materials (Sewell, W.R., Willand, J. and Craver, B.N., Surg., Gynecol. Obstet., 100, 483, 1955) Histological grading of tissue reaction Characteristic

Weighting factor

Width of zone Overall cell density Neutrophils Giant cells Lymphocytes-plasma cells Macrophages Eosinophils Fibroblasts-fibrocytes Edema Necrosis

5 3 6 2 1 1 1 2 2 3

A sample score is completed as follows: Parameter Zone Cell density Marcophages Giant cells Fibroblasts Total score

Grade

Weighting factor

Score

2 2 2 1 2

5 3 1 2 1

10 6 2 2 2 22

Surface degradation and microenvironmental outcomes

603

weighting factors of several inflammatory parameters are subjective and should be revised based on more than 30 years of published suture research information. They concluded that the development of an adapted, reliable, and reproducible scoring system is needed to assess tissue reactions to modern biomaterials. In their study of tissue reaction to ten suture materials (Silk, TevdekÕ, TicronÕ, EthibondÕ, ProleneÕ, plain catgut, chromic catgut, DexonÕ, VicrylÕ, and MaxonÕ) in male Wistar rats (in the abdominal facial layer) using conventional histologic criteria, Smit found that the data obtained failed to demonstrate any systematic differences in tissue reaction to these ten suture materials at seven days post-implantation. They suggested the large tissue reaction from surgical trauma during the first seven days post surgery masked any differences in tissue reaction caused by different biomaterials. Their data suggest that a longer period (>7±14 days) should be used to assess any difference in the tissue reaction to different biomaterials. However, they argued that longer periods of assessment were not clinically relevant because the basis of physiological wound healing is established within the first seven to fourteen days postoperation, the same period in which surgical complications including wound infection and adhesions develop. Overall, they came to the conclusion that the best means of achieving normal wound healing without complications is to minimize surgical trauma through careful tissue handling and surgical manipulation rather than the selection of biocompatible suture materials for wound closure. The general picture of cellular response to both absorbable and nonabsorbable biomaterials involves various types of cells. For the first few days, neutrophils predominate and are eventually replaced by a macrophage population with a varying content of eosinophils, lymphocytes, and plasma cells. Blood vessels gradually infiltrate the wound area and fibroblasts proliferate. After a few weeks, the reaction subsides and a band of fibrous connective tissues forms, encapsulating the implanted biomaterial. The width of the band and the number of surrounding macrophages and giant cells are related to the severity of the tissue reaction. It is generally true that the more inert the implanted biomaterial, the narrower the fibrous connective tissue band and the less the cellular reaction. However, individual variation among animals and patients does exist. Also, if the biomaterial is of a multifilament nature, the filaments are widely separated and surrounded by invading cells. Postlethwait used absorbable PGA suture fibers (DexonÕ) as a model to examine tissue reaction in rabbit by a modified Sewell's criterion (Postlethwait, 1970). As indicated in Table 21.4, the reaction reached a peak around the end of the first week and then subsided to as low as a 10.0 grade by six months. Polymorphonuclear leukocytes and lymphocytes were rare; giant cells were more common. An unmodified Sewell grading system was used to evaluate the tissue response of absorbable Vicryl braided fibers in rats (Craig et al., 1975). In that study, Craig et al.'s mean tissue score was different from Postlethwait's data. Craig's data indicated that the highest tissue response scores occur around

No. animals

12 16 10 13 13 11 10 16 7

Interval

3 days 7 days 14 days 28 days 42 days 2 months 4 months 6 months 8 months

PGA, polyglycolic acid.

12 16 10 13 13 11 10 16 7

3 days 7 days 14 days 28 days 42 days 2 months 4 months 6 months 8 months

a

No. animals

Interval

Grade of tissue reaction

38 63 46 47 56 39 21 31 15

No. sutures

54 77 46 59 33 34 11 6 0

No. sutures

34.1 38.9 42.2 42.0 44.6 46.6 21.3 12.1 18.8

Av

Chromic 3-0 Range 14.50 23.53 24.58 26.63 21.75 28.71 9.46 9.33 9.48

35.6 43.2 31.9 29.4 24.2 25.1 10.2 9.0 ö

Av

22.53 28.60 15.50 20.40 10.30 6.34 9.15 9.90 ö

PGA 0a Range

38 60 40 52 52 51 52 69 32

No. sutures

47 71 31 55 48 26 4 2 2

No. sutures

29.54 28.63 20.47 20.44 15.39 15.40 15.27 5.23 15.32

Silk Range

23.59 26.59 19.45 14.44 15.31 6.34 9.11 9.11 9.11

PGA 3-0 Range

41.1 43.4 31.9 26.2 28.2 24.0 18.5 13.1 19.2

Av

35.7 42.0 29.1 26.5 24.7 21.3 10.5 10.0 10.0

Av

28 60 37 38 62 39 52 65 32

No. sutures

36 48 36 54 36 39 35 38 24

No. sutures

12.53 23.58 19.47 11.43 15.45 15.34 11.25 5.25 5.28

Dacron Range

23.54 28.53 19.69 20.67 26.65 23.82 9.48 9.53 9.37

Chromic 0 Range

36.5 40.9 31.8 25.9 25.2 22.4 17.5 15.2 16.7

Av

36.9 37.9 43.0 41.8 42.2 48.4 26.4 21.1 16.6

Av

Table 21.4 Sewell grading of tissue reactions of a variety of sutures in New Zealand white rabbits (Postlethwait, R.W., Arch. Surg., 101, 489, 1970)

Surface degradation and microenvironmental outcomes

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21.5 Tissue reaction and absorption profiles of 4/0 polyglactin 910 and polyglycolic acid sutures at 42 days post-implantation in rat. (a) polyglactin 910 is extensively infiltrated by cells at  85; (b) Polyglactin 910. The cell population consists of an admixture of fibroblasts and macrophages at  212; (c) polyglycolic acid. The cell population occupies the spaces between the filaments at  85; (d) same as (c) except at  212. In both (b) and (d), the cross-sections of the filaments are eosinophilic and extensively fissured (Craig et al., 1975 with permission).

two months after implantation, while Postlethwait's results showed this to occur at seven days. The discrepancy could be attributed to the animal model as well as surgical technique used. The peak tissue reaction at two months observed by Craig et al. appeared to be consistent with the biodegradation and absorption of the biomaterial. This is because absorbable biomaterials normally reach their peak tissue reaction when they are biodegraded at the peak rate. Craig also compared the tissue response to PGA with that to Vicryl sutures and found that they were both quantitatively and qualitatively similar (Craig et al., 1975). Figure 21.5 details such a comparison. Those biodegradable polymers that absorbed at longer durations exhibit similar tissue response characteristics as PGA and Vicryl. For example, Bezwada et al. reported the median tissue reaction scores of Monocryl suture fibers in gluteal muscles of female Long-Evan rats (Bezwada et al., 1995). As observed in PGA and Vicryl cases, polymorphonuclear leukocytes, macrophages, fibroblasts, lymphocytes, and plasma cells and occasional foreign-body giant cells were

606

Surfaces and interfaces for biomaterials

present at the interface between Monocryl and its implant site, and these inflammatory cell densities gradually reduced during the 119 day implantation period as Monocryl was gradually absorbed. Different physical forms of biodegradable polymers appear not to alter the basic in vivo biodegradation phenomena, but could alter the time of appearance of each histological response. For example, Zhao et al. examined the histological response of Vicryl and Resolut membranes implanted subcutaneously in rats (Zhao et al., 2000). At the end of 21 days post-implantation, the classical fibroconnective capsule formation around Vicryl membrane became thinner but with pronounced giant cells, the membrane remained intact. At the end of the same period, Resolut membrane, however, was perforated and surrounded by many foreign-body giant cells. In the case of shifting to a different time of appearance of histological response due to different physical forms of biodegradable devices, typical examples are biodegradable orthopedic devices like pins and screws. Sirlin et al. reported that 80% of the 175 PDS in solid pin form was found to be completely absorbed by 24 months in a total 59 osteochondral allografts of the knee (Sirlin et al., 2001), but PDS suture fibers are completely absorbed between 6±8 months (Chu et al., 1997). The tissue response of synthetic absorbable polymers in a variety of tissues has also been reported (Bergman et al., 1971; Salthouse et al., 1977; Zhao et al., 2000, Sirlin et al., 2001). Bergman et al. examined tissue reactions to PGA sutures in three tissues; the intestine, urinary bladder, and abdominal muscle of rabbits. In the intestine, the suture was found to be embedded in fibroblast-rich granulation tissue with multinuclear giant cells of a foreign-body type, but without a fibrous capsule at the end of one month. No fibrous granulomas, as observed with catgut suture, were found around the PGA sutures. The histologic changes in the muscle and bladder were essentially the same as those in the intestine. In the three ophthalmic tissues, Salthouse et al. showed polyglactin 910 sutures induced similar cellular populations to chromic collagen sutures before 21 days as shown in Table 21.5. Except in the cornea, giant cells along with fibroblasts were found with polyglactin 910 after 21 days. Giant cells, however, were not observed with either catgut or collagen sutures in the cornea, sclera, and ocular muscle of rabbits. However, a further study raises concern about using PGA sutures in ophthalmology (Yamanaka, 1992). Yamanaka in Japan reported two clinical cases of DexonÕ suture-induced chorioretinitis in squint surgery and he suggested that PGA sutures may act as strong antigens that could lead to inflammation of delayed hypersensitivity. The lesions included macula edema, granulomatous uveitis (caused by macrophages) and capillary occulusions. Based on a mouse animal model, Yamanaka found that macrophage blocking with carrageenan and an incomplete adjuvant were necessary to evoke the delayed hypersensitivity of ophthalmologic tissue to PGA sutures. There were a large number of lymphocytes and monocytes surrounding the PGA while substantially

Slight P++ M++ Slight P++ M+ Slight P+ M++ Minimal P+ M+ Minimal P+ M+ Slight P+ M++

Minimal P+ M++ Minimal P+ M++ Slight P+ M+

Chromic surgical gut, size 7-0; cornea

Polyglactin 910, size 7-0; cornea

Polyglactin 910; ocular muscle

Polyglactin 910; sclera

Chromic collagen; ocular muscle

Chromic collagen, sclera

Chromic collagen, size 7-0; cornea

Chromic surgical gut; ocular muscle

Chromic surgical gut; sclera

7 days

Suture and site

Moderate M++

Slight M+

Minimal M+

Moderate M++

Slight M++

Slight M+

Moderate M++

Moderate P+ M++ Moderate M++

14 days

Slight M++ G+ Slight M++ G+, F+

Minimal M++

Slight M++ F Slight M++ F+

Moderate absorbing M+++ Moderate absorbing M+++ Moderate M++ F+ Slight M+

21 days

Slight M++ F+, G+ Slight M++ F+, G+

Slight M++

Slight M++ F+ Slight M++ F+

Slight M++ F+ Moderate M++ L+, F+ Slight M+

Slight M++

28 days

Nearly absorbed M++ F+ Absorbed; a few residual cells Slight M++ F+ Slight M++ F+, G+

Absorbed

Nearly absorbed M+ Slight M+ F+, L+ Slight M++ M+, F+ Absorbed M+

35 days

Slight M+ F+ Absorbed

Absorbed

Nearly absorbed

Absorbed

Absorbed M+

42 days

Absorbed

Only slight evidence of site A few residual cells Absorbed

60 days

Table 21.5 Histologic response of absorbable sutures in opthalmologic tissues (Salthouse, T.N., Maltalaga, B.F., and Wykoff, M.H., `Comparative tissue response to six suture materials in rabbit cornea, sclera, abnd ocular muscle', Am. J. Ophthalmol., 84(2), 224, 1977)

608

Surfaces and interfaces for biomaterials

fewer lymphocytes were noted in the control group of animals. Thus, Yamanaka concluded that PGA suture is not completely safe in ophthalmologic use. The degree of the tissue reaction depends largely on the chemical nature and physical configuration of the biodegradable biomaterial. For example, protein based biomaterials, such as silk and collagen, elicit far more marked tissue reactions than those of a synthetic nature like aliphatic polyesters. The physical configuration of biomaterials is a unique factor that closely relates to the level of inflammatory response, particularly at the late stage. A typical example of the physical configuration factor is multifilament vs. monofilament based biomaterials. It is well known that multifilament sutures induce more inflammatory response than monofilament sutures of an identical chemical structure. The reason behind this difference in tissue reaction, even though they are chemically identical, is that multifilament biomaterials could provide many interstitial free spaces among the filaments for inflammatory cells to penetrate and to elicit an inflammatory reaction upon in contact with individual filament surfaces. Monofilament-based biomaterials do not have these interstitial free spaces and hence elicit a lower level of inflammatory reaction. Figure 21.5 illustrates such infiltration of inflammatory cells into multifilament polyglycolide (DexonÕ) and poly(glycolide-co-lactide) (VicrylÕ) suture fibers after 42 days post-implantation in rats. A mixture of macrophages and fibroblasts occupy the interstitial space within those biodegradable multifilament sutures. Obviously, due to the high surface area multilament fibers provide, the tissue reaction would increase correspondingly. The physical form of sutures, ie., multifilament vs. monofilament, also affected the level of fibrosis in repaired vessel walls. Gersak (1990) reported that monofilament sutures resulted in less fibrosis in vessel walls compared with multifilament sutures. Along with the fibrosis of arterial walls, Gersak also found a higher level of persistent calcification in vessel walls closed with absorbable PDSÕ sutures than with non-absorbable ProleneÕ sutures. The major link between the surface area and shape of an implant and its tissue reaction can be demonstrated by some landmark work done on suture fibers and surgical meshes (Matlaga et al., 1976, 1980; Salthouse, 1985). In those studies, using rat gluteal muscle as the model, Matlaga et al. observed that macrophages closely adhered to the Vicryl absorbable suture and appeared to surround the filaments by extension of thin lamellapodia as early as one day post-implantation. This type of cell-implant surface interaction was not generally observed with non-biodegradable suture materials. Fibroblasts and collagen were evident at five days and reached their highest concentration at 28 days. Absorption was evident from 35 to 53 days and was accompanied by the appearance of large numbers of mitochondria in the close vicinity of the plasma membranes of macrophages and foreign-body giant cells adjacent to the filaments. Such evidence of increased metabolic activity was believed to be related to the metabolism of the hydrolytic degradation products of Vicryl, glycolate, and lactate. It was the authors' opinion

Surface degradation and microenvironmental outcomes

609

that this related to the timing for the peak appearance of absorption of absorbable biomaterials. Matlaga et al. also found that triangular-shaped implants elicited the highest enzyme (lysosomal acid phosphatase) activity, followed by pentagon shape and then circular shape with the lowest enzyme activity level (Matlaga et al., 1976). This relationship applies to a wide range of biomaterials from polyolefins, polyurethanes to silicone rubber. In Salthouse's reported study (Salthouse, 1985), Vicryl, Dacron, Nylon, and polypropylene and Vicryl suture fibers and Vicryl surgical mesh fabrics were implanted in rat gluteal muscle and tissue reactions were evaluated by SEM examination of stained tissue blocks and enzyme histochemistry. The most interesting observation was that the shape and surface of an implant can profoundly affect macrophage behavior which, in turn, could affect how fast biodegradable implants would be destroyed and how wound healing would proceed. First, there is firm attachment of macrophages onto the implant surface, regardless of the biodegradable or non-biodegradable nature of the biomaterial. The adhered macrophage shows prominent rough endoplasmic reticulum and Golgi zones with pinocytotic vesicles at the cell surface. It is believed such a firm adherence onto an implant surface would activate macrophages and is responsible for the biodegradation of that implant as well as the inflammatory response associated with that implant. This is because macrophages can control fibroblast function, such as collagen synthesis, during wound healing, via macrophage secretion of lactate which is known to stimulate collagen synthesis (Hunt et al., 1978). Such an increase in lactate level was found at polylactide mesh implantation sites (Salthouse, 1985) and was associated with an increased rate of wound healing associated with polylactide mesh. Second, different distinctive macrophage behavior in terms of numbers of adhered cells and enzyme activity (glucose-6-phosphate dehydrogenase) was found over a period of 90 days at a smooth vs. rough surface of the same biomaterial. For example, a surfaceroughen Teflon rod (1 mm diameter) had a macrophage population many times higher at its interface than the same rod with a smooth surface. Similar findings were also reported in cultured macrophages in which active macrophages and giant cells were at the interface with an implant having rough surface at three months, while no macrophages were found on the surface of a smooth sample after four weeks (Rich et al., 1981). Therefore, it appears prudent to design implants having smooth and well rounded physical finishes for better tissue biocompatibility. As described above, Salthouse and Matlaga suggested that the close adhesion of macrophages onto a biomaterial surface like Vicryl, Dacron, Nylon, and polypropylene surface bears a direct relationship to the observed in vivo biodegradation properties. Two published studies that provide a good demonstration of the relationship between the surface degradation of biomaterials and their microenvironmental outcomes in vivo are those of Zhong et al. on absorbable Vicryl suture placed inside a poly(methylmethacrylate) chamber implanted in

610

Surfaces and interfaces for biomaterials

21.6 Inflammatory response of tissue to implanted sutures. (a) Response to Prolene and PDS; (b) response to maxon; () Prolene implanted in fascia; (l) PDS implanted in fascia; (n) PDS implanted in peritoneum (Metz et al., 1989 with permission).

rats (Zhong et al., 1993) and of Metz et al. in a study of two absorbable monofilament sutures (PDS and Maxon) in rabbits (Metz et al., 1989). Metz et al. examined in vivo tissue reactivity and surface morphology related degradation of both absorbable MaxonÕ and PDSÕ monofilament sutures in rabbit peritoneum and rectus fascia for intervals of two to 70 days for the purpose of identifying whether the complications of abdominal surgery (i.e., intraperitoneal inflammation and adhesion formation) relate to suture-induced tissue reaction (Metz et al., 1989). As shown in Fig. 21.6 the tissue inflammatory response (defined as the number of inflammatory cells/mm2) to both MaxonÕ and PDSÕ was similar. Inflammation reached a maximum at four days post-implantation and then subsided to preoperative levels by day 21. The inflammatory response was more intense in peritoneum than in fascia. The similar in vivo inflammatory reaction to PDS and Maxon, however, was not reflected in their observed surface morphology. Figure 21.7 illustrates the SEM of both MaxonÕ and PDSÕ sutures retrieved from rabbit peritoneum or fascia at 21 and 35 days post-implantation. Contrary to Zhong et al.'s reported difference in the surface morphology of Vicryl between controlled in vivo and in vitro degradation environments, absorbable monofilament MaxonÕ sutures in both fascia and peritoneum tissues for 35 days (Fig. 21.7) exhibited the formation of surface cracks and subsequent peeling of the outermost layer like onion skins, while PDS did not show any visible surface morphological change (Fig. 21.7) for the same tissue and duration. This similar observed tissue inflammatory response but a different surface morphological appearance for the implanted absorbable biomaterials

Surface degradation and microenvironmental outcomes

611

21.7 Scanning electron micrographs of implanted sutures in New Zealand white rabbits. (a) PDS suture in peritoneum at 21 days,  175; (b) PDS suture in peritoneum at 21 days,  43; (c) Maxon suture in fascia at 35 days,  11; (d) Maxon suture in peritoneum at 35 days,  21 (Metz et al., 1989 with permission).

suggests that although both Maxon and PDS have the same monofilament structure, the chemical structure difference between these two biomaterials and their subsequent different fiber manufacturing processes are believed to be the basis for the significant observed difference in surface morphological response. The Metz et al. findings could also suggest that the microenvironment due to the in vivo inflammatory response appears to depend more on the physical configuration than the chemical structure of the biomaterials, even though the chemical difference would lead to a different surface morphology upon biodegradation in vivo. Therefore, the surface characteristics of a biomaterial appears to be more important than the bulk of the same biomaterial to control the

612

Surfaces and interfaces for biomaterials

level of in vivo inflammatory response, while the bulk characteristic of a biomaterial would be more critical for controlling the surface morphological appearance upon biodegradation in vivo. The onion-peeling like surface morphology observed in Maxon in rabbit tissues is also similar to the surface morphology observed in the in vitro PBS

21.8 SEM micrographs of Maxon sutures after 42 days in vitro hydrolytic degradation in PBS at 37 ëC. (a) 0.2 Mrad -irradiation before immersion; (b) 2.0 Mrad -irradiation before immersion (Chu et al., 1995 with permission).

Surface degradation and microenvironmental outcomes

613

environment shown in Fig. 21.8. In conjunction with these surface morphologic observations, MaxonÕ suture became very fragile and disintegrated easily after 14 days, but PDSÕ sutures retained structural integrity for at least 35 days implantation. The similarity of the unique surface degradation-induced morphology of Maxon between the in vitro PBS microenvironment (Fig. 21.8) and the in vivo tissue environment (Fig. 21.7) may suggest that enzymes may not play a major role in the unique surface morphology upon degradation since no enzymes are present in the in vitro PBS microenvironment. The interactions between surface degradation property of synthetic biodegradable polymers and their microenvironment have been the foundation for expanded applications of this class of synthetic biodegradable biomaterials beyond their conventional areas (e.g. wound closure and healing). One exciting non-conventional use of such interactions between biodegradable polymers and their surrounding host is tissue engineering, which has become a major research domain because of the advantage of the lack of permanent foreign-body reaction induced by the presence of biomaterial scaffolds. For example, in the study conducted by Partridge et al., human bone marrow cells transduced with a specific gene were seeded onto PLGA (copolymer of PGA and PLA) scaffolds and the cells grew on the biodegradable scaffold (Partridge et al., 2002). The bone marrow cells proliferated further and differentiated on the PLGA scaffold after implantation in mice. Extensive cell growth was observed and there was physical evidence for the formation of cartilage and bone tissue by the gene transduced in the bone marrow cells. The scaffolds were biodegraded by normal hydrolysis and the site was replaced by natural tissue without the long-term complications that are found with foreign implants. Thomson et al. used PLGA foam scaffolds for guided tissue fabrication and they concluded that PLGA foam scaffolds were biocompatible and efficient conductors for tissue growth (Thomson et al., 1999). There was no evidence of inflammatory cells or signs of adverse tissue reaction to the PLGA foam, and there was no evidence of tissue damage from the degradation products. Bare PLA and PGA surface is not an ideal surface for cell growth and differentiation. As a result, there have been many attempts to modify the surface of these aliphatic biodegradable polymers. One of the most frequently used methods is to treat the biodegradable polymer surface with RGD peptides (ArgGly-Asp) to make the surface more cell friendly. A recent example of such a surface modification for tissue engineering is the immobilization of poly(Llysine) attached RGD peptide sequence onto the surface of PLA and PGLA substrates (Yang et al., 2001; Quirk et al., 2001). The RGD peptide-modified PLA and PGLA surface enhanced osteoblast cell attachment, spreading and proliferation, and the level of enhancement depended on the solution concentration of the RGD peptides. At an optimal RGD solution concentration used for PLA surface coating, the osteoblast cell adhesion and spreading were found to be comparable to tissue culture plate serum controls.

614

Surfaces and interfaces for biomaterials

21.5 Conclusion The biodegradation phenomenon of synthetic absorbable polymers is controlled by the chemical nature of the polymers, their physical forms and surrounding microenvironment. As a result, a wide range of biodegradation property could be observed from a wide range of commercially available synthetic absorbable polymers. In vivo biodegradation of these commercially important absorbable biomaterials appears to activate macrophage during inflammatory response. Such a macrophage activation is the pivotal step that subsequently leads to the cascade of biochemical and biological events of wound healing as well as regeneration of tissues. Hence, controlling the biodegradation property of biodegradable biomaterials should provide a unique means to control macrophage-induced biochemical activities. Such a capability would be a powerful strategy for using biodegradable substrates for engineering tissues.

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poly(glycolic acid) suture material as a function of applied strain', Biomaterials, 5, 365±368. Mizuma, K, Lee, P C and Howard, J M (1977), The disintegration of surgical sutures in exposure to pancreatic juice, Ann. Surg., 186, 718±722. Niemi, M and Sylven, B (1968), `On the chemical pathology of interstitial fluid', Acta. Path. Microbil. Scand., 72, 205±219. Partridge K, Yang X, Clarke N, et al. (2002), `Adenoviral BMP-2 Gene Transfer in Mesenchymal Stem Cells: in vitro and in vivo Bone Formation on Biodegradable Polymer Scaffolds', Biochemical and Biophysical Research Communications. March, 292(1), 144±152. Persson, M, Bilgrav, K, Jensen, L and Gottrap, F (1986), `Enzymatic wound cleaning and absorbable sutures: An experimental study on varidase and dexon sutures', Eur. Surg. Res., 18, 122±128. Postlethwait, R W (1970), `Polyglycolic acid surgical suture', Arch. Surg., 101, 489±494. Potts, J E (1978), `Biodegradation', in Jellinek, H H G (ed.), Aspects of Degradation and Stabilization of Polymers, Elsevier, Amsterdam, pp. 617±657. Poulsen, J, Kristensen, Y N, Brygger, H E and Delikaris, P (1983), `Treatment of infected surgical wounds with varidase', Acta. Chir. Scand., 149, 245±248. Pratt, L, Chu, A, Kim, J, Hsu, A and Chu, C C (1993), `The effect of electrolytes on the in vitro hydrolytic degradation of synthetic biodegradable polymers: Mechanical properties, thermodynamics and molecular modeling', J. Polym Sci. Chem edn, 31, 1759±1769. Quirk, R A, Chan, W C, Davies, M C, Tendler, D S and Shakesheff, K M, (2001), `Poly(L-lysine)-GRGDS as a biomimetic surface modifier for poly(lactic acid)', Biomaterials, 22, 865-872. Rich, A and Harris, A K (1981), `Anomalous preferences of cultured macrophages for hydrophobic and roughened substrata', J. Cell Sci., 5, 1±7. Salthouse, T (1976), `Cellular enzyme activity at the polymer ± tissue interface: A review', J. Biomed. Mater. Res., 10, 197±229. Salthouse, T N (1985), `Some aspects of macrophage behavior at the implant interface', J. Biomed. Mater. Res., 18(4), 395±401. Salthouse, T N and Matlaga, B F (1976), `Polyglactin 910 suture absorption and the role of cellular enzymes', Surg. Gynecol. Obstet., 142, 544±550. Salthouse, T N, Matlaga, B F and Wykoff, M H (1977), `Comparative tissue response to six suture materials in rabbit cornea, sclera, and ocular muscle', Am. J. Ophthalmol., 84, 224±233. Sanz M, Zabalegui I, Villa A and Sicilia A, (1997), `Guided tissue regeneration in human class II furcations and interproximal infrabony defects after using a bioabsorbable membrane barrier', Internl. J. Periodontics and Restorative Dentistry, 17(6), 563± 573. Schmitt, E E and Polistina, R A (1967), `Surgical sutures', U.S. patent 3,297,033, American Cyanamid. Schnabel, W (1981), Polymer Degradation: Principles and Practical Applications, Hanser International, Wien, p. 112. Shalaby, S W (1994), Biomedical Polymers: Designed-to-Degrade Systems, Hanser Publishers, New York. Sewell, W R, Willand, J and Craver, B N (1955), Surg. Gynecol. Obstet., 100, 483. Sirlin, C B, Doutin, R D, Brossmann, J, Pathria, M N, Convery, F R, Bugbee, W and

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Resnick, D, (2001), `Polydioxanone biodegradable pins in the knee: MR Imaging', Amer. J. Radiology, 176, 83±90. Smit, I B, Witte, E, Brand, R and Trimbos, J B, (1991), `Tissue reaction to suture materials revisited: Is there argument to change our views', Eur. Sug. Res., 23, 347. Thomson R, Mikos A, et al., (1999), `Guided Tissue Fabrication from periosteum using preformed biodegradable polymer scaffolds', Biomaterials. 20(21), 2007±2018. Tonetti M S, Cortellini, P, Suvan, J E, et al., (1998), `Generalizability of the added benefits of guided tissue regeneration in the treatment of deep intrabony defects. Evaluation in a multi-center randomized controlled clinical trial', J Periodontology, 69(11), 1183±1192. Vert, M, Feijen, J, Albertsson, A, Scott, G and Chiellini, E (1992), Biodegradable Polymers and Plastics, Royal Soc of Chem., Cambridge, England. Williams, D F (1979), `Some observations on the role of cellular enzymes in the in vivo degradation of polymers', ASTM Spec. Tech. Publ., 684, 61±75. Williams, D F (1982), `Biodegradation of surgical polymers', J. Mater. Sci., 17, 1233. Williams, D F and Chu, C C (1984), `The effects of enzymes and gamma irradiation on the tensile strength and morphology of poly(p-dioxanone) fibers', J. Appl. Polym. Sci., 29, 1865±1877. Williams, D F and Mort, E (1977), `Enzyme-accelerated hydrolysis of polyglycolic acid', J. Bioeng. 1, 231±238. Yamanaka, T (1992), `Chorioretinitis caused by synthetic absorbable sutures', Lens and Eye Toxicity Res. 9 (3,4), 559. Yang, X B, Roach, H I, Clarke, N M P, Howdle, S M, Quirk, R, Shakesheff, K M and Oreffo, R O C, (2001), `Human osteoprogenitor growth and differentiation on synthetic biodegradable structures after surface modification', Bone, 29(6), 523± 531. Zaikov, G E (1985), `Quantitative aspects of polymer degradation in the living body', J. Macromol. Sci. ± Rev. Macromol. Chem. Phys., C25, 551±597. Zhao, S J, Pinholt, E M, Madsen, J E and Donath, K (2000), `Histological evaluation of different biodegradable and non-biodegradable membranes implanted subcutaneously in rats', J. Cranio-Maxillofacial Surg., 28, 116±122. Zhong, S P, Doherty, P J and Williams, D F (1993), `The degradation of glycolic acid/ lactic acid copolymer in vivo', Clinical Materials, 14, 145. Zhong, S P, Doherty, P J and Williams, D F (1994), `A preliminary study on the free radical degradation of glycolic acid/lactic acid copolymer', Plastics, Rubber and Composites Processing and Application, 21, 89±97.

22

Microbial biofilms and clinical implants

M M I L L A R , Barts and the London School of Medicine and Dentistry

22.1 Introduction The areas covered in this chapter include mechanisms by which clinical implants facilitate biofilm formation and infection, the epidemiology, risk factors and microbiology of clinical implant infections, consequences of biofilm formation, costs, diagnosis, treatment, and prevention of implant-associated infections, and potential areas for future research. Information resources and references are also included.

22.1.1 Mechanisms by which clinical implants facilitate microbial colonisation and biofilm formation The formation of biofilms on clinical implants can be demonstrated by a variety of microscopic techniques (see for example Peters et al., 1981). The physicochemical interactions leading to biofilm formation on implants have been reviewed in other chapters of this book. Clinical implants may become contaminated with microorganisms prior to placement of the implant, or may become colonised either from spread of bacteria from a contiguous site, from a distant site through the bloodstream or through a natural barrier to infection that has been breached by the implant. There are a number of potential mechanisms by which implants can predispose to the establishment of biofilm. Implants have a considerable impact on host resistance to infection. Implants can damage tissues, compromise the local vascular supply, provide protected niches for microbial proliferation, compromise local immunity (such as through complement depletion), and also provide a surface for microbial attachment and for biofilm formation. Biofilms on clinical implants almost always form within a complex matrix of host-derived macromolecules. Binding of specific host-derived macromolecules to the implant surface may facilitate microbial attachment through specific binding receptors. Frequently, the implant surface has been conditioned by host factors prior to microbial attachment. Biofilm formation is a prerequisite for the development of a clinical implant infection.

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22.2 Epidemiology and costs of infection associated with clinical implants Clinical implants are perhaps the major risk factor for hospital-acquired infection. It is estimated that 10% of hospital patients will develop a hospitalacquired infection and a large proportion of these infections arise in association with a clinical implant, such as a urethral catheter, tracheal tube or vascular catheter (Emmerson et al., 1996). There are estimated to be more than 5000 deaths/year directly attributable to hospital-acquired infection (Department of Health, 1995). These infections also add greatly to the cost of healthcare and may lead to considerable prolongation of a patient's stay in hospital (see Plowman et al., 1999). More than 50% of hospital-acquired blood stream infections in the UK are device-related ± most frequently a central vascular catheter (Waghorn, 1994; Fletcher and Bodenham, 1999; NINSS, 2002). The central vascular catheter infection rate varies from 4±15 episodes/1000 days of central line use, depending upon the patient population. The cost of a central vascular catheter associated infection can be many thousands of pounds per episode depending upon the virulence of the infecting agent. A recent report from the Chief Medical Officer (Department of Health, 2003) outlines the potential benefits of improved prevention of hospital-acquired infection. There are seven areas for action identified ranging from improved surveillance to research. Action point two covers some of the good practice points associated with prevention of implant-associated infections. In a study using a US database of over 9000 patients ventilated on an intensive care unit in 1998/9, 9% developed ventilator-associated pneumonia (VAP). Each episode of VAP increased hospital charges by US$40,000 (Rello et al., 2002). VAP is much more likely to result from infection with antibioticresistant bacteria, and so is much more difficult to treat and has a high associated mortality (Chastre and Fallon, 2002). The risk of pneumonia is 3±21 times greater when the patient is intubated (Celis et al., 1988; Chastre and Fallon, 2002). The infection rates for prosthetic joint implants in the UK vary from <1 to >10% depending upon the unit, risk factors and the type of procedure (NINSS, 2001). Lower rates have been reported (Phillips et al., 2003). The costs (excluding loss of earnings) of a prosthetic joint infection have been estimated at around US$50,000, and in addition to the pain and disability there is also a significant associated mortality (Lentino, 2003). The risk of a clinical implant associated infection is determined by factors related to the patient, implant design and materials used in the composition, insertion site, procedure, and operating environment, management of the implant after insertion, and the way in which the implant is used so, for example, some infusates such as intralipid may be important risk factors for central vascular catheter infection (Freeman et al., 1990) and a recent report suggests

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that catecholamine inotropes may promote biofilm formation by Staphylococcus epidermidis (Lyte et al., 2003). Patient-related factors include age (Langley et al., 2001), general health, immune status, and compliance of staff or patients with preventive strategies. It is most important that staff involved in the insertion of implants are appropriately skilled and trained. Many serious clinical implant infections such as central vascular catheter associated infections are a consequence of individuals failing to comply with protocols for the safe management of the implant. Implant design and materials are important determinants of the risk of microbial colonisation, biofilm formation and associated infection. Aspects of design include the ease of implantation, required degree of operator skill, extent of tissue damage at the time and following implantation, and susceptibility of the implant to contamination following implantation. Materials used may also be important determinants of the risk of infection. Surface charge, hydrophobicity, and roughness are all determinants of microbial attachment (Fletcher and Loeb, 1979) and biofilm formation. Some biomaterials produce toxic substances such as plasticisers which may cause local tissue damage, increasing susceptibility to microbial colonisation. The conditions of use of the biomaterial, for example microbial/surface interactions under conditions of flow and in different ionic environments may also be an important determinant of biofilm formation (Millar et al., 2001). Biomaterials may be modified by the incorporation of antimicrobials or by surface modification in an attempt to reduce the risk of bacterial colonisation and infection. There is evidence that these strategies can significantly reduce the frequency of microbial infection associated with implants. Design of connections may be an important determinant of the risk of infection for implants which have an external component that requires frequent manipulations such as a central vascular or urethral catheter. For further details of prevention of biofilm formation on implants see section below.

22.3 Microbiology of clinical implant infections Almost all infections associated with clinical implants are caused by bacteria or fungi. The bacteria associated with implant infections include Gram positive bacteria such as staphylococci, enterococci, corynebacteria and Gram negative bacteria such as enterobacteriaceae, Pseudomonas spp. and Acinetobacter spp. Candida spp. (Douglas, 2003) is the most frequent group of yeasts associated with implant infections. Obligate intracellular parasites such as viruses and also protozoa and worms are rarely associated with clinical implant infections. Amoebae may cause infection associated with contact lenses (Beattie et al., 2003). In the natural world and at sites of microbial colonisation such as the oral cavity, biofilms usually consist of polymicrobial consortia. Biofilms associated

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with clinical implants may be polymicrobial but more often consist of single strains of bacteria. Identification of the cause of an implant infection can have an important impact on patient management. The most frequent bacteria associated with implant infections are staphylococci particularly Staphylococcus epidermidis. The strains of Staphylococcus epidermidis which colonise implants, may be derived from the skin of healthcare staff or the patient. Infections caused by Staphylococcus aureus are much more likely to result in serious adverse consequences for the patient, so that treatment of a Staphylococcus aureus infection associated with an implant almost always involves removal of the implant, combined with an extended duration of treatment. Implant-related infections acquired in hospital are frequently caused by antibiotic-resistant bacteria. These include infections with methicillin-resistant Staphylococcus aureus (MRSA). Patient outcomes with antibiotic-resistant infections tend to be much worse than those caused by antibiotic-sensitive bacteria.

22.4 Molecular mechanisms underlying biofilm formation A considerable amount of knowledge has accumulated over the last 20 years on the mechanisms by which microorganisms adhere to human living tissues at sites of microbial colonisation (for an overview see Donlan, 2002). There are human host factors which inhibit biofilm formation by some bacteria. It has recently been reported that low concentrations of the host defence protein lactoferrin promotes twitching motility in Pseudomonas aeruginosa, and that this is associated with a reduction in biofilm formation (Singh et al., 2002). More recently the mechanisms underlying microbial interactions with implant surfaces have begun to be elucidated (Harshey, 2003). Microorganisms can adhere directly to implants or bind to host-derived macromolecules. The initial process that leads to attachment of a microorganism to a surface can involve non-specific physicochemical interactions (van der Waals forces, hydrophobic and charge interactions) or binding of specific microbial components to the implant surface (Veenstra et al., 1996), or to host-derived macromolecules attached to the polymer surface such as fibronectin (Menzies, 2003). Binding of bacteria directly to the implant surface is probably most important immediately prior to and at the time of implant insertion. Specific host protein binding molecules have been identified on the surface of some bacterial species (Herrmann et al., 1988). The mechanisms of surface-sensing by bacteria and the potential to subvert early interactions between bacteria and implants as a preventive strategy is reviewed by Lejeune (2003). Microbial surface characteristics such as flagellae and pili can be important determinants of biofilm development (Landini and Zehnder, 2002; Klausen et al., 2003). Bile has been reported as an important determinant of biofilm formation on

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gallstones by Salmonella typhi (Prouty et al., 2002). Strains of Pseudomonas aeruginosa that colonise the lungs of patients with cystic fibrosis may produce large amounts of alginate, and it was assumed that over-production of alginate is a colonisation factor in these patients. Recent studies suggest that alginate overproduction is not required for colonisation by Pseudomonas aeruginosa (Wozniak et al., 2003). Following attachment, individual cells multiply and form microcolonies. These microcolonies aggregate as they grow. The microbial surface growth matures from an amorphous layer to develop a biofilm architecture. This architecture consists of bacterial cells in a hydrated matrix. Convection currents move fluids through the bulk of the biofilm through interconnected channels. The processes leading to channel formation are poorly understood but may include variations in growth rates resulting from accumulation of inhibitors, and substrate limitation (Chang et al., 2003). The genetics and regulation of biofilm development, have been studied particularly in Escherichia coli (Shembri et al., 2003), and Pseudomonas aeruginosa (see for example O'Toole, 2003), and increasingly in staphylococci (Dobinsky et al., 2003; Caiazza and O'Toole, 2003; von Eiff et al., 2002). There is increasing evidence that the expression of specific genes is required for biofilm formation. The expression of these genes may be regulated by growth conditions and cell density (see Wagner et al., 2003). In some bacteria such as Pseudomonas aeruginosa the accumulation of signalling metabolites (as cell density increases) directly modifies gene expression (quorum sensing) (Whitehead et al., 2001; Schuster et al., 2003). N-acyl homoserine lactone is the best known of these quorum signalling metabolites. In addition to modifying bacterial cell gene transcripton these signalling molecules may also modify gene expression in eukaryotic cells (Smith and Iglewski, 2003). Eukaryotic cell responses to bacterial quorum-sensing molecules may be an important determinant of the outcome of host defence/biofilm interactions (Mathesius et al., 2003). The evolution of quorum signalling and its significance for microbial ecology remains to be elucidated (Manefield and Turner, 2002; Bornberg-Bauer and Weiner, 2002). Interference with these cell signalling mechanisms could be used as a mechanism for controlling microbial colonisation of implants (Hentzer and Givskov, 2003). In Staphylococcus epidermidis and Staphylococcus aureus a cluster of genes controlling production of an intercellular adhesin and biofilm formation has been described (McKenney et al., 1998; Crampton et al., 1999). Accumulation of the ica encoded polysaccharide intercellular adhesin (PIA) protects the bacterial cells from antimicrobial factors produced by or administered to the host. A number of other factors may also be important in biofilm formation by staphylococci (for an overview see Gotz, 2002; and also von Eiff et al., 2002). Coaggregation of bacteria of different species may have a role in the development of multispecies biofilm (Rickard et al., 2003). The majority of implant infections involve a single microbial species. There is little information

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on the significance of interactions between bacteria of different species or genera in the pathogenesis or persistence of implant infections. It has been suggested that inter-species interactions may be important determinants of colonisation at carrier sites (see for example Lina et al., 2003).

22.5 Determinants of biofilm antibiotic resistance Bacterial biofilms tend to show much higher levels of resistance to antimicrobial agents than planktonic cells (Hoyle and Costerton, 1991). There are many potential mechanisms to explain this increased resistance. Growth phase and growth rate are key determinants of antimicrobial susceptibility (Cozens et al., 1986). Slow growing bacteria are much less susceptible to the bactericidal activity of antimicrobials such as beta-lactams. A proportion of the microorganisms in a biofilm are nutrient deprived and grow very slowly (Anderl et al., 2003). Binding of antimicrobials to extracellular matrix macromolecules and dead bacteria, protection of viable bacteria deep in the matrix by inactivating enzyme activity from surrounding bacteria (Ciofu, 2003) and changes in ionic gradients leading to changes in antimicrobial diffusion may also contribute some protection to biofilm bacteria. Alterations in microbial physiology associated with the biofilm mode of growth probably also contribute to reducing antimicrobial susceptibility. A recent report suggests that binding of antimicrobials to periplasmic glucans may be an important mechanism of resistance of biofilm bacteria to some antibiotics in Pseudomonas aeruginosa (Mah et al., 2003). Microelectrode analysis of oxygen concentrations in biofilm layers suggests that only the bacteria at the air/biofilm interface have sufficient oxygen for growth, and so it may be that oxygen utilisation and poor diffusion is at least as important as the diffusion of antimicrobial agents in determining antimicrobial resistance in Pseudomonas aeruginosa biofilms (Walters et al., 2003). A genetic locus has been described in Pseudomonas aeruginosa which is important in regulating phenotype (slow-growing, small colony variants), propensity to biofilm formation and antibiotic resistance (Drenkard and Ausubel, 2002).

22.6 Consequences of biofilm formation on clinical implants In clinical practice, biofilm formation on a clinical implant generally becomes apparent only when the patient develops signs or symptoms of infection, or biofilm formation leads to compromise of the function of the implant. The natural history of biofilm development is poorly understood. So, for example, we do not know how frequently biofilms are controlled by a local immune response without the development of clinical signs or symptoms of disease. A breakdown in good infection control practice with respect to management of a central vascular catheter can result in a clinical presentation of infection days or

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weeks later. Determinants of variation in individual susceptibility to implant infection have not been elucidated. Bacterial colonisation of an implant can lead to calcification (Sung et al., 1993) so, for example, colonisation of a urinary catheter with Proteus mirabilis, causes alkalinisation of urine (a bacterial urease releases ammonia from urea) and precipitation of insoluble calcium salts. Precipitated salts encrust the implant and may form bladder stones. Reports implicating nanobacteria in biomineralisation (Kajander and Ciftcioglu, 1998) have not been confirmed (Cisar et al., 2000). A biofilm inside a vascular or urinary catheter may lead to blockage of the catheter lumen. A local inflammatory response around a joint prosthesis may lead to breakdown in the bone surrounding the prosthesis, with loosening of the prosthesis (Perdreau-Remington et al., 1996). Many of the bacterial species most frequently associated with implant infection such as Staphylococcus epidermidis are of low virulence, so that the clinical presentation may be chronic, insidious and non-specific (Jansen et al., 1991). An inflammatory reaction to biofilm formation by a low-grade pathogen such as Staphylococcus epidermidis may be an important factor in the development of a fibrous capsular contracture around breast implants (Pajkos et al., 2003). Infections with Staphylococcus epidermidis, even in severely immunocompromised patients, are rarely fatal. On the other hand, infections caused by Staphylococcus aureus and some Gram-negative species such as Klebsiella spp. can have an acute onset with a significant associated mortality.

22.7 Clinical implant infection 22.7.1 Diagnosis Clinical implant infections can be easily diagnosed when a patient with an implant and few other risk factors for infection develops signs and symptoms of infection associated with inflammation at the site of the implant. A more frequent scenario is the insidious development of non-specific signs and symptoms of infection. The difficulty of diagnosing infection associated with an implant is illustrated by the very high rate at which central vascular catheters are inappropriately removed for suspected infection (Farr, 1999). Diagnostic difficulties arise because of the difficulty in sampling the surface of the implant. Diagnostic difficulties underlie the ongoing uncertainty concerning the role of biofilm formation in the chronic changes found around some types of implant, so for example capsular contractures which develop around breast implants may be a consequence of biofilm formation at the implant surface (Pajkos et al., 2003). In some populations of patients, diagnosis based on the culture of bacteria is also compromised by use of antibiotics. Clinicians may give a trial of antibiotic therapy for a suspected implant-associated infection at an inaccessible site such

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as a replacement joint. Antibiotics may also be administered for the prevention of infection in immunocompromised patients. Attempts to diagnose infection at the implant site, even with invasive sampling methods such as biopsy, may fail to lead to the culture of an aetiological agent. Diagnostic methods based on DNA detection may lead to improvements in diagnostic accuracy when samples can be obtained from within or surrounding the implant. A method of diagnosis of central venous catheter associated infections based on eubacterial DNA amplification has recently been reported (Warwick et al., 2004). Newer radiological approaches such as positron emission tomography (PET) (Schiesser et al., 2003) and the use of labelled antimicrobial peptides (Lupetti et al., 2003) are also being investigated. Detection of specific immune responses may also have a role in diagnosis (Selan et al., 2002).

22.7.2 Treatment Implants also provide niches for microbial proliferation which are protected from host defence factors and administered antimicrobials. For these reasons, and others alluded to above, microorganisms growing in implant-associated biofilms are much more resistant to antimicrobial agents and host defence factors than planktonic cells, and successful management of infection associated with a clinical implant may require removal of the implant. Some types of infection will respond to treatment with systemic antibiotics. Central vascular catheter colonisation associated with Staphylococcus epidermidis bacteraemia frequently responds to the infusion of antibiotics through the catheter lumen. The optimum antimicrobial strategy for the eradication of bacteria from central vascular catheters remains to be determined, so for example the optimum duration of exposure of biofilm bacteria to antimicrobials has not been ascertained (Ley et al., 1996) and the potential benefit of combinations requires further work (Peck et al., 2003). Outcome is far better with early removal of the central vascular catheter when associated with fungal colonisation and infection. Effective treatment of Staphylococcus aureus bloodstream infections associated with colonisation of a central vascular catheter requires early removal of the catheter and a prolonged duration of systemic antibiotic treatment (Lowy, 1998). Antibiotics, even if from the same class, may vary in their activity against biofilm bacteria for reasons that remain to be elucidated (Di Bonaventura et al., 2004). Diagnostic microbiology laboratories usually use planktonic cells for the antimicrobial susceptibility tests that are used to inform patient management. Methods have been reported for the susceptibility testing of clinical isolates growing in biofilms, but these methods are relatively time-consuming and expensive, and have not yet been widely incorporated into routine laboratory practice (Knobloch et al., 2002; Labthavikul et al., 2003). The Calgary Biofilm Device allows a large number of antimicrobial agents and combinations of agents

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to be compared for activity against biofilms in a 96 well plate (Ceri et al., 1999; 2001). A 96 well plate method which relies on measurement of formazan salt reduction has been reported for antifungal susceptibility testing in Candida albicans (Ramage et al., 2001). A novel mouse model designed for the detection of the activity of antimicrobials against biofilms in living animals has been reported (Kadurugamuwa et al., 2003a,b). This model relies on the detection of bioluminescent reporters by biophotonic imaging, and has the potential to allow the assessment of the anti-biofilm activity of both new and existing agents in vivo. There is considerable potential for the expanding knowledge of the molecular mechanisms underlying biofilm formation to lead to new strategies for the control or treatment of biofilms on clinical implants (see, for example, Hentzer et al., 2003). RNA III inhibiting peptide (RIP) has been shown to inhibit biofilm formation by Staphylococcus aureus, and when combined with antibiotics, inhibits biofilm formation by Staphylococcus aureus on Dacron grafts in Wistar rats (Giacometti et al., 2003). RIP may also show synergistic activity against Staphylococcus epidermidis (Balaban et al., 2003). A potential problem with use of more effective anti-biofilm strategies is that many of the bacteria that contribute to our health live at mucosal surfaces in biofilm communities, and these populations may also be disrupted by the more effective anti-biofilm treatments.

22.8 Prevention of biofilm formation on clinical implants Inevitably, prevention of infection involves prevention of biofilm formation. The potential for studies of biofilms to inform the development of strategies for the control of implant-associated infections has been reviewed recently (Costerton et al., 2003). Effective preventive strategies may require novel implant designs and surface coatings. A number of methods have been reported for the study of microbial interaction with surfaces. These methods allow the interaction of bacteria with different surface materials to be compared. More recently described methods such as the parallel plate flow chamber (Bos et al., 1999) and methods using the modified Robbins device (Millar et al., 2001) allow study of microbial surface interactions under conditions of flow. Indirect methods of measuring microbial attachment to surfaces have also been described (Poortinga et al., 1999). Animal models for studying biofilm infections of specific types of implant have also been described (Sheehan et al., 2004). Silver ions are rapidly bactericidal against a wide range of bacteria, and silver preparations have been used to prevent infection of areas of damaged skin such as those resulting from burns. Increasingly silver is being used and proposed as an antimicrobial constituent of biomaterials (Sabbuba et al., 2002; Multanen et al., 2002; Corral et al., 2003). The extent to which increasing use of silver will lead to an increase in silver resistance in bacteria is not known (Silver, 2003).

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Changes in human behaviour may also result in a reduction in the frequency of implant-related infection so, for example, infections can be prevented through the development and implementation of evidence-based good practice guidelines (see for example Kilbride et al., 2003a,b). Some of the preventive strategies for urethral catheters, intravascular catheters, tracheal tubes and prosthetic joint infections are outlined below. Carriage of Staphylococcus aureus is a risk factor for subsequent Staphylococcus aureus infection and at least 50% of Staphylococcus aureus infections are endogenous, so eradication of the carriage of Staphylococcus aureus has been used as a successful strategy to control infection in some groups of patients (Luzar et al., 1990). A number of strategies for reducing microbial attachment and growth by implant surface modification have been described. Some of which are mentioned below.

22.8.1 Urethral catheters Urethral catheterisation is the most significant risk factor for hospital-acquired urinary tract infection, which itself is one of the most frequent infections acquired by patients in hospital (Sedor and Mulholland, 1999). Urethral catheters may be used in the short-term management of a patient, such as for drainage of the bladder, prior to a pelvic procedure. Long-term catheters are used for patients with chronic conditions resulting in urinary retention. Ensuring that staff who undertake urethral catheterisation are adequately trained, and minimising the number of disconnections of the catheter (sealed units) are important measures that limit the frequency of infection (Warren, 1997). It is possible to prevent infection associated with short-term catheterisation by aseptic catheter insertion, careful management of the drainage system and appropriate use of antimicrobials (Saint and Lipsky, 1999), but reliable strategies for the long-term prevention of urethral catheter-associated infection remain to be found. Considerable work is on-going to identify novel materials, designs or coatings that might limit biofilm formation (Beiko et al., 2003; Stickler et al., 2002). Antimicrobial agents that are used for systemic infection such as ciprofloxacin have been included in coatings, but these have the drawback of selecting for resistance to agents of great value in the treatment of patients with serious infection. There is some data suggesting that silver alloy catheters can reduce the frequency of catheter-associated urinary tract infection (Saint et al., 2000a). A recent report describes the prevention of biofilm formation on Foley (urethral) catheters in a laboratory model by the inflation of the retention balloon with disinfectant solution (triclosan) (Stickler et al., 2003). Avoidance of urethral catheterisation would prevent biofilm formation and infection. Surprisingly, physicians may not be aware that patients have an indwelling urethral catheter ± particularly when the catheter has been put in inappropriately! (Saint et al., 2000b). Guidelines for the prevention of urethral

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catheter-associated infection have been published by the UK National Institute for Clinical Excellence (www.NICE.org.uk).

22.8.2 Central vascular catheters Central vascular catheter (CVC) associated infections are some of the most serious and expensive of hospital-acquired infections. Guidelines on the prevention of CVC-associated infections have been published by the UK National Institute for Clinical Excellence (www.NICE.org.uk). As with all implants, CVCs should be implanted by an experienced and trained operator. The optimum choice of insertion method, site, and catheter will depend on assessment of the individual. Risk factors and preventive strategies will vary in different patient populations, so prevention of CVC-associated infection in neonates (Kilbride et al., 2003a,b) may require a different emphasis from that required for adults in an intensive care unit, or undergoing treatment for cancer. Heparin bonded catheters and those impregnated with antimicrobial agents reduce the rate of infection associated with short-term CVCs (Dobson, 2003; Shah et al., 2002). In a recent report of a meta-analysis (Walder et al., 2002) it was concluded that antibiotic and chlorhexidine-silver sulfadiazine coatings of CVCs are anti-infective for a week or so. There was evidence of lack of effect for silver-impregnated collagen cuffs.

22.8.3 Tracheal tubes and ventilator associated pneumonia Most of the effective strategies for the prevention of VAP rely on strict adherence to infection control protocols such as dedicated use of disposable suction catheters, and the physical management of the airways and ventilation circuits. Education of staff in current best practice recommendations can reduce the rates of VAP (Zack et al., 2002). Use of topical and systemic antibiotics with the aim of reducing upper respiratory colonisation with enteric Gram negative bacteria (selective digestive decontamination) has also been shown to prevent VAP, but may be associated with an increased risk of selection antibiotic resistant bacteria. There are a number of recent reviews of current strategies for the prevention of VAP (for example Craven et al., 2002; Rello et al., 2002; Collard et al., 2003). Research in animals has shown that coating of the inside of the tracheal tube with an antibacterial agent may reduce biofilm formation (Berra et al., 2003). Entrainment of biofilm in the ventilated gases as a consequence of turbulent flow through the tracheal tube may be an important factor in the aetiology of VAP (Inglis et al., 1989). Avoiding the use of a tracheal tube through the use of non-invasive positive pressure ventilation has the potential to reduce considerably the incidence of VAP (Sinuff and Cook, 2003).

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22.8.4 Prosthetic joints Data from the UK National Nosocomial Infection Surveillance Scheme shows considerable variation in the proportion of hip and knee prostheses that become infected (PHLS, 1999), even when the results take account the patient risk status. This suggests that much could be done to reduce infections associated with these procedures. Prevention of prosthetic joint infections requires careful attention to the general health of the patient pre-operatively, optimising intraoperative management and operative conditions, and post-operative care (Douglas et al., 2001). Most prosthetic joint procedures are carried out in operating theatres with a laminar air flow system with a filtered air supply, which reduces the chance of airborne bacteria contaminating the operation site (Norden, 1985, 1991). Prophylactic antibiotics further reduce the risk of infection (Ann and Friedman, 1996). Surgical technique and the types of materials and designs used in prosthetic joints are probably also important determinants of the risk of biofilm formation and infection.

22.9 Further research As the use and diversity of clinical implants grows, it is likely that the prevention of implant-associated infections will present an increasing challenge. In many countries, rates of implant infection are increasingly viewed as an indication of the quality of patient care, and it is likely that there will be an increasing emphasis on mandatory reporting of infection rates (see for example House of Commons, 2000), and also an increasing emphasis on prevention. The development of implants with a reduced frequency of infective complications is a major challenge to manufacturers. This challenge will inevitably involve multidisciplinary research and experts from a wide range of fields. Research will require collaboration between clinicians, paramedical staff and basic scientists. Areas of research range from clinical trials of novel implants, determination of best practice for the insertion and management of implants, ways of ensuring compliance with best practice, selection of materials used in manufacture, implant design, and methods and materials used for implant surface conditioning.

22.10 Information resources The Center for Biofilm Engineering (http://www.erc.montana.edu). The American Society for Microbiology (ASM). The ASM produce A Manual of Biofilm Related Exercises. The UK Society for General Microbiology. Organised symposia and meetings which cover all aspects of Biofilm Microbiology. Methods in Enzymology has volumes 310, 336 and 337 covering many aspects of

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biofilm science. The Biofilm Club (http://www.biofilmclub.co.uk). The Biofilm Journal. This on-line journal has links to a range of biofilm-related information sources.

22.11 References Anderl J N, Zahller J, Roe F and Stewart P S (2003), `Role of nutrient limitation and stationary-phase existence in Klebsiella pneumoniae biofilm resistance to ampicillin and ciprofloxacin', Antimicrob Agents Chemother, 47 (4), 1251±1256. Ann Y H and Friedman R J (1996), `Prevention of sepsis in total joint arthroplasty', J Hosp Infect, 33 (2), 93±108. Balaban N et al. (2003), `Use of the quorum-sensing inhibitor RNAIII-inhibiting peptide to prevent biofilm formation in vivo by drug-resistant Staphylococcus epidermidis', J Infect Dis, 187 (4), 625±630. Beattie T K et al. (2003), `Enhanced attachment of acanthamoeba to extended-wear silicone hydrogel contact lenses: a new risk factor for infection?', Ophthalmology, 110 (4), 765±771. Beiko D T et al. (2003), `Biomaterials in urology', Curr Urol Rep, 4 (1), 51±55. Berra L et al. (2003), `New approaches for the prevention of airway infection in ventilated patients. Lessons learned from laboratory animal studies at the National Institutes of Health', Minerva Anestesiol, 69 (5), 342±347. Bornberg-Bauer E and Weiner III J (2002), `Quorum sensing in context: out of molecular biology and into microbial ecology', Microbiology, 148, 3762±3765. Bos R, van der Mei H C and Busscher H J (1999), `Psycho-chemistry of initial microbial adhesive interactions ± its mechanisms and methods for study', FEMS Microbiol Rev, 23, 179±230. Caiazza N C and O'Toole G A (2003), `Alpha-toxin is required for biofilm formation by Staphylococcus aureus', J Bacteriol, 185 (10), 3214±3217. Celis R et al. (1988), `Nosocomial pneumonia. A multivariate analysis of risk and prognosis', Chest, 93, 318±324. Ceri H et al. (1999), `The Calgary Biofilm Device: new technology for rapid determination of antibiotic susceptibilities of bacterial biofilms', J Clin Microbiol, 37 (6), 1771±1776. Ceri H et al. (2001), `The MBEC Assay System: multiple equivalent biofilms for antibiotic and biocide susceptibility testing', Methods Enzymol, 337, 377±385. Chang I, Gilbert E S, Eliashberg N and Keasling J D (2003), `A three-dimensional, stochastic simulation of biofilm growth transport-related factors that affect structure', Microbiology, 149 (Pt 10), 2859±2871. Chastre J and Fallon J Y (2002), `Ventilator-associated pneumonia', Am J Respir Crit Care Med, 165, 867±903. Ciofu O (2003), `Pseudomonas aeruginosa chromosomal beta-lactamase in patients with cystic fibrosis and chronic lung infection. Mechanism of antibiotic resistance and target of the humoral immune response', Apmis, Suppl, 116, 1±47. Cisar J O et al. (2000), `An alternative interpretation of nanobacteria-induced biomineralization', Proc Natl Acad Sci USA, 97 (21), 11511±11515. Collard H R et al. (2003), `Prevention of ventilator-associated pneumonia: an evidence-

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Giacometti A et al. (2003), `RNA III inhibiting peptide inhibits in vivo film formation by drugresistant Staphylococcus aureus', Antimicrob Agents Chemother, 47 (6), 1979±1983. Gotz F (2002), `Staphylococcus and biofilms', Mol Microbiol, 43 (6), 1367±1378. Harshey R M (2003), `Bacterial motility on a surface: many ways to a common goal', Annu Rev Microbiol, 57, 249±273. Hentzer M and Givskov M (2003), `Pharmacological inhibition of quorum sensing for the treatment of chronic bacterial infections', J Clin Invest, 112 (9), 1300±1307. Hentzer M et al. (2003), `Attenuation of Pseudomonas aeruginosa virulence by quorum sensing inhibitors', EMBO J, 22 (15), 3803±3815. Herrmann M et al. (1988), `Fibronectin, fibrinogen, and laminin act as mediators of adherence of clinical staphylococcal isolates to foreign material', J Infect Dis, 158, 693±701. House of Commons (2000), 42nd Report of the Public Accounts Committee. The Management and Control of Hospital Acquired Infection in Acute NHS Trusts in England. Hoyle B D and Costerton W J (1991), `Bacterial resistance to antibiotics: the role of biofilms', Prog Drug Res, 37, 91±105. Inglis T J J , Millar M R, Gareth-Jones J and Robinson D A (1989). `Tracheal tube biofilm as a source of bacterial contamination of the ventilated lung', J Clin Microbiol 27: 2014±2018. Jansen B et al. (1991), `Late onset endophthalmitis associated with intraocular lens: a case of molecularly proved S epidermidis aetiology', Br J Ophthalmol, 75, 440±441. Kadurugamuwa J L et al. (2003a), `Direct continuous method for monitoring biofilm infection in mouse model', Infect Immun, 71 (2), 882±890. Kadurugamuwa J L et al. (2003b), `Rapid direct method for monitoring antibiotics in a mouse model of bacterial biofilm infection', Antimicrob Agents Chemother, 47 (10), 3130±3137. Kajander E O and Ciftcioglu N (1998), `Nanobacteria: an alternative mechanism for pathogenic intra- and extracellular calcification and stone formation', Proc Natl Acad Sci USA, 95, 8274±8279. Kilbride H W, Wirtschafter D D, Powers R J and Sheehan M B (2003a), `Implementation of evidence-based potentially better practices to decrease nosocomial infections', Pediatrics, 111 (4), 519±533. Kilbride H W et al. (2003b), `Evaluation and development of potentially better practices to prevent neonatal nosocomial bacteremia', Pediatrics, 111 (4 Pt 2): e504±e518. Klausen M et al. (2003), `Biofilm formation by Pseudomonas aeruginosa wild type, flagella and type IV pili mutants', Mol Microbiol, 48 (6), 1511±1524. Knobloch J K, Von Osten H, Horstkotte M A, Rohde H and Mack D (2002), `Minimal attachment killing (MAK): a versatile method for susceptibility testing of attached biofilm-positive and -negative Staphylococcus epidermidis', Med Microbiol Immunol (Berl), 191 (2), 107±114. Labthavikul P, Petersen P J and Bradford P A (2003), `In vitro activity of tigecycline against Staphylococcus epidermidis growing in an adherent-cell biofilm model', Antimicrob Agents Chemother, 47 (12), 3967±3969. Landini P and Zehnder A J (2002), `The global regulatory hns gene negatively affects adhesion to solid surfaces by anaerobically grown Escherichia coli by modulating expression of flagellar genes and lipopolysacchar production', J Bacteriol, 184 (6), 1522±1529.

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Langley J M et al. (2001), `Unique epidemiology of nosocomial urinary tract infection in children', Am J Infect Control, 29 (2), 94±98. Lejeune P (2003), `Contamination of abiotic surfaces: what a colonizing bacterium sees and how to blur it', Trends in Microbiology, 11 (4), 179±184. Lentino J R (2003), `Prosthetic joint infections: bane of orthopedists, challenge for infectious disease specialists', Clin Infect Dis, 36 (9), 1157±1161. Ley Bm Jalil N, Foote A, Wilson M and Millar M R (1996), `Bolus or infusion teicoplanin for intravascular catheter associated infections in immunocompromised patients?', J Antimicrob Chemotherap, 38, 1091±1095. Lina G, Boutite F, Tristan A, Bes M, Etienne J and Vandenesch F (2003), `Bacterial competition for human nasal cavity colonization: role of staphylococcal agr alleles', Appl Environ Microbiol, 69, 18±23. Lowy F D (1998), `Staphylococcus aureus infections', N Engl J Med, 339 (8), 520±532. Lupetti A et al. (2003), `Radiolabelled antimicrobial peptides for infection detection', Lancet Infect Dis, 3 (4), 223±229. Luzar M A et al. (1990), `Staphylococcus aureus nasal carriage and infection in patients on continuous ambulatory peritoneal dialysis', N Engl J Med, 322 (8), 505±509. Lyte M et al. (2003), `Stimulation of Staphylococcus epidermidis growth and biofilm formation by catecholamine inotropes', Lancet, 361 (9352), 130±135. Mah T F et al. (2003), `A genetic basis for Pseudomonas aeruginosa biofilm antibiotic resistance', Nature, 426 (6964), 306±310. Manefield M and Turner S L (2002), `Quorum sensing in context: out of molecular biology and into microbial ecology', Microbiol, 148, 3762±3764. Mathesius U et al. (2003), `Extensive and specific responses of a eukaryote to bacterial quorum-sensing signals', Proc Natl Acad Sci USA, 100 (3), 1444±1449. McKenney D et al. (1998), `The ica locus of Staphylococcus epidermidis encodes production of the capsular polysaccharide/adhesin', Infect Immun, 66 (10), 4711± 4720. Menzies B E (2003), `The role of fibronectin binding proteins in the pathogenesis of Staphylococcus aureus infections', Curr Opin Infect Dis, 16 (3), 225±229. Millar M R et al. (2001), `Use of a continuous culture system linked to a modified Robbins device or flow cell to study attachment of bacteria to surfaces', in Doyle R J, Methods in Enzymology, California, Academic Press, 43±62. Multanen M et al. (2002), `Biocompatibility, encrustation and biodegradation of ofloxacine and silver nitrate coated poly-L-lactic acid stents in rabbit urethra', Urol Res, 30 (4), 227±232. Norden C W (1985), `Prevention of bone and joint infections', Am J Med, 78 (6B), 229± 232. Norden C W (1991), `Antibiotic prophylaxis in orthopedic surgery', Rev Infect Dis, 13 (Suppl 10), S82±S846. Nosocomial Infection National Surveillance Scheme (2001), Surveillance of Surgical Site Infection in English Hospitals 1997±2001. Public Health Laboratory Service. Nosocomial Infection National Surveillance Scheme (2002), Surveillance of HospitalAcquired Bacteraemia in English Hospitals 1997±2002. Public Health Laboratory Service. O'Toole G A (2003), `To build a biofilm', J Bacteriol, 185 (9), 2687±2689. Pajkos A et al. (2003), `Detection of subclinical infection in significant breast implant capsules', Plast Reconstr Surg, 111 (5), 1605±1611.

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Peck K R et al. (2003), `Antimicrobials as potential adjunctive agents in the treatment of biofilm infection with Staphylococcus epidermidis', Chemotherapy, 49 (4), 189±193. Perdreau-Remington F et al. (1996), `A four-year prospective study on microbial ecology of explanted prosthetic hips in 52 patients with ``aseptic'' prosthetic joint loosening', Eur J Clin Microbiol Infect Dis, 15, 160±165. Peters G et al. (1981), `Microbial colonization of prosthetic devices. II. Scanning elecron microscopy of naturally infected intravenous catheters', Zentralbl Bakteriol Mikrobiol Hyg B, 173, 293±299. Phillips C B et al. (2003), `Incidence rates of dislocation, pulmonary embolism, and deep infection during the first six months after elective total hip replacement', Bone Joint Surg Am, 85-A (1), 20±26. Plowman R et al. (1999). `The socio-economic burden of hospital acquired infection', Public Health Laboratory Service. Poortinga A T, Bos R and Busscher H J (1999), `Measurement of change transfer during bacterial adhesion to an indium tin oxide surface in a parallel plate flow chamber', J Microbiol Methods, 38, 183±189. Prouty A M, Schwesinger W H and Gunn J S (2002), `Biofilm formation and interaction with the surfaces of gallstones by Salmonella spp', Infect Immun, 70 (5), 2640±2649. Public Health Laboratory Service, Surveillance of surgical site infection in English hospitals 1997±1999, London, PHLS. Ramage, G, Vande Walle K, Wickes B L and Lopez-Ribot J L (2001), `Standardized method for in vitro antifungal susceptibility testing of Candida albicans biofilms', Antimicrob Agents Chemother, 45 (9), 2475±2479. Rello J et al. (2002), `Epidemiology and outcomes of ventilator-associated pneumonia in a large US database', Chest, 122, 2115±2121. Rickard A H et al. (2003), `Bacterial coaggregation: an integral process in the development of multi-species biofilms', Trends Microbiol, 11 (2), 94±100. Sabbuba N, Hughes G and Stickler D J (2002), `The migration of Proteus mirabilis and other urinary tract pathogens over Foley catheters', BJU Int, 89 (1), 55±60. Saint S and Lipsky B A (1999), `Preventing catheter-related bacteriuria: should we? Can we? How?', Arch Intern Med, 159 (8), 800±808. Saint S et al. (2000a), `The potential clinical and economic benefits of silver alloy urinary catheters in preventing urinary tract infection', Arch Intern Med, 160 (17), 2670±2675. Saint S et al. (2000b), `Are physicians aware of which of their patients have indwelling urinary catheters?', Am J Med, 109 (6), 476±480. Schiesser M et al. (2003), `Detection of metallic implant-associated infections with FDG PET in patients with trauma: correlation with microbiologic results', Radiology, 226 (2), 391±398. Schuster M et al. (2003), `Identification, timing and signal specificity of Pseudomonas aeruginosa quorum-controlled genes: a transcriptome analysis', J Bacteriol, 185 (7), 2066±2079. Sedor J and Mulholland S G (1999), `Hospital-acquired urinary tract infections associated with the indwelling catheter', Urol Clin North Am, 26 (4), 821±828. Selan L et al. (2002), `Diagnosis of vascular graft infections with antibodies against staphylococcal slime antigens', Lancet, 359 (9324), 2166±2168. Shah P S, Ng E and Sinha A K (2002), `Heparin for prolonging peripheral intravenous catheter use in neonates (Cochrane Review)', Cochrane Database Syst Rev, CD002774.

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Sheehan E et al. (2004), `Adhesion of Staphylococcus to orthopaedic metals, an in vivo study', J Orthop Res, 22 (1), 39±43. Shembri M A, Kjaergaard K and Klemm P (2003), `Global gene expression in Escherichia coli biofilms', Mol Microbiol, 48 (1), 253±267. Silver S (2003), `Bacterial silver resistance: molecular biology and uses and misuses of silver compounds', FEMS Microbiol Rev, 27 (2±3), 341±353. Singh P K, Parsek M R, Greenberg E P and Welsh M J (2002), `A component of innate immunity prevents bacterial biofilm development', Nature, 417, 552±555. Sinuff T and Cook D J (2003), `Health technology assessment in the ICU: noninvasive positive pressure ventilation for acute respiratory failure', J Crit Care, 18 (1), 59±67. Smith R S and Iglewski B H (2003), `P. aeruginosa quorum-sensing systems and virulence', Curr Opin Microbiol, 6 (1), 56±60. Stickler D J, Evans A, Morris N and Hughes G (2002), `Strategies for the control of catheter encrustation', Int J Antimicrob, 19 (6), 499±506. Stickler D J et al. (2003), `Control of encrustation and blockage of Foley catheters', Lancet, 361 (9367), 1435±1437. Sung J Y, Leung J W, Shaffer E A, Lam K and Costerton J W (1993), `Bacterial biofilm, brown pigment stone and blockage of billiary stents', J Gastroenterol Hepatol, 8 (1), 28±34. Veenstra G J et al. (1996), `Ultrastructural organization and regulation of a biomaterial adhesin of Staphylococcus epidermidis', J Bacteriol, 178, 537±541. von Eiff C et al. (2002), `Pathogenesis of infections due to coagulase-negative staphylococci', Lancet Infectious Diseases, 2, 677±685. Waghorn D J (1994), `Intravascular device associated systemic infections: a 2-year analysis of cases in a district general hospital', Journal of Hospital Infection, 28, 91±101. Wagner V E et al. (2003), `Microarray analysis of Pseudomonas aeruginosa quorumsensing regulons: effects of growth phase and environment', J Bacteriol, 185 (7), 2080±2895. Walder B et al. (2002), `Prevention of bloodstream infections with central venous catheters treated with anti-infective agents depends on catheter type and insertion time: evidence from a meta-analysis', Infect Control Hosp Epidemiol, 23 (12), 748± 756. Walters MC III, Roe F, Bugnicourt A, Franklin MJ and Stewart P S (2003), `Contributions of antibiotic penetration, oxygen limitation, and low metabolic activity to tolerance of Pseudomonas aeruginosa biofilms of ciprofloxacin and tobramycin', Antimicrob Agents Chemother, 47 (1), 317±323. Warren J W (1997), `Catheter-associated urinary tract infections', Infect Dis Clin North Am, 11 (3), 609±622. Warwick S, Wilks M, Hennessy E, Powell-Tuck J, Small M, Sharp J and Millar M R (2004), `Diagnosis of central vascular catheter associated bacterial infection using quantitative 16S rDNA detection', J Clin Microbiol, 42 (4), 1402±1408. Whitehead N A, Barnard A M, Slater H, Simpson N J and Salmond G P (2001), `Quorumsensing in Gram-negative bacteria', FEMS Microbiol Rev, 25, 365±404. Wozniak D J et al. (2003), `Alginate is not a significant component of the extracellular polysaccharide matrix of PA14 and PA01 Pseudomonas aeruginosa biofilms', Proc Natl Acad Sci USA, 100 (13), 7907±7912. Zack J E et al. (2002), `Effect of an education program aimed at reducing the occurrence of ventilator-associated pneumonia', Crit Car Med, 30 (11), 2407±2412.

23

Extracellular matrix molecules in vascular tissue engineering C M K I E L T Y , D V B A X , N H O D S O N and M J S H E R R A T T , Wellcome Trust Centre for Cell-Matrix Research, UK

23.1 Introduction Vascular disease is the major cause of death in Western society. Arterial bypass operations using autogenous grafts are a common treatment, with more than 600,000 procedures conducted per annum in the USA alone. Coronary artery bypass operations are used to relieve the symptoms of angina and lifethreatening cardiovascular disease, but there is a significant morbidity and cost associated with bypass operations. Many patients do not have suitable veins (usually internal mammary artery or saphenous vein), there is a compliance mismatch between veins and arteries, and thrombosis and infection are also serious problems. Thus, there is a strong demand for alternative patent vascular prostheses based on biomaterials. Technological advances in tissue engineering have resulted in the generation of vascular substitutes that increasingly resemble natural vessels.1 However, many problems with these vascular prostheses remain, including EC detachment during flow leading to thrombosis, lack of smooth muscle cell (SMC) proliferation and limited deposition of vascular extracellular matrix (ECM). Thus, there is a need to incorporate appropriate vascular ECM molecules and other biological signals such as growth factors into prostheses. ECM molecules and growth factors can encourage migration of endogenous cells into grafts from neighbouring vessel walls, direct cell adhesion, proliferation and ECM deposition within the scaffold, and regulate vascular cell phenotype, for example, by controlled delivery and release of vascular growth factors such as TGF 1, PDGF-BB and VEGF. We are developing small diameter vascular grafts based on a porous elastomeric scaffold incorporating ECM molecules, with vascular SMC seeded within the scaffold using biodegradable threedimensional ECM-based gels. In order to highlight the matrix molecules that will appropriately influence vascular cells within grafts, it is necessary to consider briefly the normal structure and composition of arteries.

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23.2 Natural blood vessels 23.2.1 Vessel wall layers and their functions Arteries comprise three major layers, the tunica intima, tunica media and tunica adventitia2 (Fig. 23.1a). The tunica intima consists of an endothelial cell (EC) monolayer which is strongly attached to a thin subendothelial ECM. It provides an anti-thrombogenic inner lining to blood vessels, a barrier to mononuclear blood cells, and a diffusion gradient to systemic cytokines and growth factors. The boundary of the tunica intima is the internal elastic lamina. The tunica media comprises alternating layers of elongated SMC with their long axis concentric to the lumen, interspersed with concentrically arranged elastic fibres, with some collagen fibre bundles and proteoglycans. It endows elastic recoil on vessel walls through the elastic fibre lamellae, and contractility though SMC. The number of elastic lamellae varies depending on the function of the vessel, with the human aortic arch having between 40±70 layers. The external elastic lamina, a thick elastic fibre layer surrounding the entire vessel circumference, separates the tunica media from the tunica adventitia. The tunica adventitia is a dense connective tissue containing smooth muscle-like myofibroblasts, abundant collagen fibres and some dispersed elastic fibres. It exerts homeostatic control on the vessel, limits extensibility, and integrates the blood vessel into the surrounding tissue.

23.2.2 Extracellular matrix components Collagens (mainly fibrillar types I, III, V) represent more than 30% of the dry weight of blood vessels.2,3 Network-forming collagens IV and VIII are major structural components of subendothelial and SMC basal laminae. Collagen VIII forms specialised and distinctive hexagonal arrays and is strongly expressed in development and following arterial injury.4 Collagen VI microfibrils are present

23.1 Structure of natural arteries. (a) Light microscopy of a cross-section through a porcine coronary artery, showing tunica intima, tunica media and tunica adventitia layers. (b) Transmission electron micrograph of the interface between media SMC and adjacent elastic fibre, highlighting the structural relationship and presence of dense plaques close to the interacting SMC surface.

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in the subendothelium, media and adventitia; these assemblies link vascular ECM molecules and cells, and contribute to the supramolecular architecture of vascular matrices. Collagen types XV and XVIII are associated with subendothelial basement membranes; their C-terminal `endostatin' fragments can be angiosuppressive or angiostimulatory and bind heparan sulphate.5,6 Elastic fibres comprise a cross-linked elastin core surrounded by an outer mantle of fibrillin microfibrils.7 Both fibrillin and elastin can interact directly with vascular SMC and EC, and strongly influence their behaviour. The expression of the soluble precursor tropoelastin is stimulated by TGF , and both tropoelastin and fibrillin-1 expression are enhanced by mechanical forces. Elastic fibres are associated with other ECM glycoproteins including fibulin-58,9 and latent TGF binding proteins (LTBPs).10 Fibulin-5 is a SMC pericellular glycoprotein that is required for normal elastic fibre formation. Over-expression of an alternatively spliced short form of versican (V3) was found to induce ordered elastic fibre formation in a vascular repair model.11 Other vascular glycoproteins include laminin in basement membranes and fibronectin which interacts with the collagens, and glycosaminoglycans that influence the polarity and migration of EC. Vitronectin, which is found within the media and adventitia, modulates cell migration, and osteopontin is sometimes found in association with medial elastic fibres; and there are also several microfibril-associated glycoproteins including MAGP-1 and latent TGF binding proteins (LTBPs).7 Proteoglycans function as structural elements of the vessel wall and modulate cell adhesion. Their biological roles arise from both their protein and carbohydrate components. They form a hydrated scaffold and are especially abundant in the subendothelium. The heparan sulphate proteoglycan, perlecan, is an intrinsic component of subendothelial and SMC basement membranes. Heparin can inhibit SMC proliferation and migration, and it also binds the growth factor FGF. Thrombomodulin has a chondroitin sulphate glycosaminoglycan (GAG) chain that enhances binding to thrombin, thereby preventing coagulation. The small leucine-rich proteoglycans decorin and biglycan are implicated in the process of elastic fibre formation,12 and regulate collagen packing.13,14 Decorin can also bind TGF growth factors. SMC-associated proteoglycans (syndecans) present in the medial SMC, together with integrin receptors, contribute to cell-matrix communication and signalling.

23.2.3 Vascular cells Endothelial cells (EC) exhibit classic cobblestone morphology in culture. In vessels, they are somewhat more elongated with their long axis parallel to blood flow, and are connected to each other by tight and gap junctions. An EC monolayer lines the lumen, provides a barrier between the blood and the vessel wall, and regulates vessel tone and leukocyte adhesion.

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Smooth muscle cells (SMC) are spindle-shaped contractile cells.15,16 Phenotypically they are highly variable in culture, usually adopting a synthetic proliferative state and forming `hills and valleys'. However, in a mature vessel wall, SMC adopt a distinctive contractile phenotype with SMC-specific cytoskeletal elements such as SM -actin, SM myosin, smoothelin and SM22. In the tunica media, SMC are surrounded by a specialised basal lamina except where they are connected to one another by gap junctions and to elastic fibres at prominent dense plaques (Fig. 23.1b). Adventitial myofibroblasts express SM -actin, exhibit an ordered cytoskeleton and have contractile properties.17,18 They deposit abundant collagen fibres and, following vascular injury, are involved in extensive adventitial remodelling. Mesenchymal stem cells (MSC): adult bone marrow stromal cells include a population of spindle-shaped adherent MSC that have SMC-like characteristics.19,20 An MSC population fated to become SMC with self-expansion abilities has been identified using a tissue-specific promoter sorting approach.21 MSC are an attractive source of SMC for vascular grafts since they can be induced to express late differentiation SMC markers following controlled growth factor treatment, and shear stress upregulates vascular ECM deposition.20,22,23

23.2.4 Vascular cell-matrix interactions EC interact strongly with their subendothelial matrix. SMC interact with elastic fibres, mainly through integrins which are heterodimeric transmembrane receptors.24 These interactions are critical for a stable endothelium and normal EC function. Following ligation of ECM molecules, the integrin -chain cytoplasmic tail interacts with talin, which in turn binds vinculin which is associated with -actinin. Thus, the vascular ECM is linked, through integrins, to the cytoskeletal framework. Integrin-mediated cell-matrix interactions regulate vascular cell survival, phenotype proliferation, migration and ECM expression and deposition. Integrins expressed by EC and SMC include integrins 1 1, 2 1, 5 1, 6 1, and v 3. Major vascular adhesion ligands include fibronectin, laminin, the elastic fibre molecules fibrillin-1, elastin, and associated molecules fibulin-5, and collagen VIII. The particular importance of elastic fibre molecules in regulating vascular cells is increasingly apparent, and below are outlined their cell adhesion characteristics. We have developed a range of recombinant elastic fibre products for seeding gels and coating vascular graft scaffolds, in order to influence SMC and EC attachment and phenotype. Fibrillin-1 is a large multidomain calcium-binding glycoprotein molecule, and the major fibrillin isoform in the adult vasculature.25 It contains a single arggly-asp (RGD) cell adhesion motif within an eight-cysteine domain that supports cell adhesion.26,27 We have shown that fibrillin-1 interacts with integrins 5 1

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and v 3 in both molecular and assembled microfibrillar forms, and supports cell adhesion, spreading, signalling and feedback upregulation of its own expression.28 Elastin interacts with SMC through the elastin binding protein, a 67-kDa protein that functions as a chaperone and facilitates the deposition of tropoelastin on the microfibril template.29 Signalling through this receptor strongly influences SMC proliferation. Fibulin-5 is a multidomain adhesive glycoprotein with a single RGD motif within a calcium-binding domain.8,9 It mediates endothelial cell adhesion, is a component of the SMC pericellular matrix, interacts with SMC via integrins 9 1, and profoundly affects elastic fibre deposition. It is a TGF-beta-inducible target gene that regulates cell growth and motility and affects protein kinase activation.30 Collagen VIII is an important vascular EC and SMC migration template.31 The 2(VIII) chain contains two RGD motifs within its collagenous sequence.

23.2.5 Vascular development During early vasculogenesis, primitive endothelial cell tubes form16,32 then a subendothelial matrix is deposited that stabilises the EC monolayer. Mesenchymal cells are recruited from surrounding tissues as SMC that first deposit vascular wall ECM then adopt a contractile phenotype. Recruitment of mesenchymal cells by EC involves growth factors such as PDGF-BB and TGF- 1.33 We have shown that direct EC contact with bone marrow-derived stem cells (MSC) profoundly influences MSC cytoskeletal organisation.20 Abundant fibrillin microfibrils are laid down in the direction of blood flow and are tightly associated with the EC subluminal surface. These microfibrils are relatively stiff elastomers,34 and their subendothelial role is to provide strong biomechanical anchorage for EC. These microfibrils also act as the template for tropoelastin deposition during tightly developmentally regulated elastic fibre formation throughout the vessel wall. They also act as a conduit for LTBPs that sequester TGF during vessel formation and repair.35 Developmental thickening of the tunica media and tunica adventitia reflects elastic fibre and collagen (mainly collagens I, III) deposition, and occurs during the first years of life. Hyaluronan (HA, an unsulphated glycosaminoglycan) is another prominent component of developing vessels. These details of how vascular ECM contributes to the development and biological properties of native arteries, are an essential backdrop for directed exploitation of key ECM molecules and supramolecular assemblies in tissue engineering of artificial grafts.

23.3 Vascular tissue engineering The concept of tissue-engineered vessels offers an attractive treatment for vascular disease.1 Several key features are required for a replacement artery to be successful. It must be infection-resistant, thrombo-resistant, biocompatible,

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leak-proof, and have appropriate mechanical properties. Also it must possess appropriate vasoactive physiological properties, and be manufactured cheaply and rapidly. Many problems remain in generating replacement vessels, including the time taken to tissue engineer arteries (seven weeks or more) which is too long for many patients, poor EC attachment, regulation of vessel tone and resistance to blood flow, thrombosis, poor survival and ECM deposition of seeded SMC within grafts, inappropriate compliance properties, and graft failure due to limited mechanical strength. In animal models, success of endothelialisation depends on cell seeding density and substrate composition. However, lack of spontaneous endothelialisation in human subjects causes thrombosis and coagulation and encourages leukocyte adhesion. Pre-seeding vascular prostheses with EC can improve the patency of small calibre grafts, but strong EC attachment to scaffolds is vital for successful endothelialisation.

23.3.1 Replacement arteries and scaffold materials Natural vessels Fresh allograft vessels are no longer used in bypass surgery due to poor patency, rejection complications, and endothelial cell sloughing. Decellularised cadaveric human arteries, another possible source of vascular conduit, have limited availability and the potential for disease transmission limits their use. Decellularised tissue-engineered arteries could potentially serve as grafts for implantation for host cell migration and colonisation.1 Several studies have quantified the effects of decellularisation on vascular matrix and mechanical properties, including cellular elimination, extracellular matrix retention, and mechanical characteristics of arteries before and after decellularisation treatments. Decellularised porcine iliac artery showed patency for at least four months, and was infiltrated by host SMC.36 A decellularisation method for native and engineered arteries was reported that maximised cellular elimination without greatly compromising mechanical integrity.37 Cell-free matrices that mimic in vivo microenvironments have also been used for studying cell-three-dimensional matrix interactions.38,39 Plating cells directly into isolated, cell-free matrices produced from cell- or tissuederived three-dimensional matrices allows rapid formation of focal adhesion, so essential structural information must be retained in these cell-free matrices. Tissue-engineered vascular substitutes Tissue engineering and the use of biomaterials have changed current approaches to vascular replacement, generating natural prostheses colonised with vascular cells in vitro that mimic natural vessels, and have the properties of growth, remodelling and repair. SMC within the scaffold are encouraged to deposit a vascular ECM that may, depending on the scaffold biodegradability properties, gradually replace the

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scaffold with an ordered functional vascular ECM rich in elastic fibres, fibrillar collagens and proteoglycans. Such tissue engineering approaches tend to build vessels in `reverse' of normal development by first establishing a strong tunica media equivalent then encouraging luminal endothelialisation.

23.3.2 Synthetic scaffolds Various synthetic materials are commonly used in engineered vascular grafts. Most, though inert, are regarded as foreign bodies by the organism and this limits their suitability over time. The synthetic materials most used are expanded polytetrafluoroethylene (ePTFE) and polyester Dacron fibres. These materials, however, are not particularly suitable for small-diameter grafts because of a tendency for thrombus induction, embolism and occlusion of the graft lumen, lack of compliance, and excessive intimal hyperplasia at anastomotic joins. Moreover, these synthetic grafts do not allow vessel remodelling or vascular physiological responsiveness, and are prone to infection. Dacron is a condensation polymer obtained from ethylene glycol and polyterephthalate, and its properties include high tensile strength and high resistance to stretching. It is markedly hydrophobic and usually does not contain basic or acidic groups. Unmodified ePTFE is hydrophobic with a low surface energy and weak electrical charge, conditions that are not optimal for EC attachment.40±42 Glow discharge and PTFE hydroxylation have improved cell adhesion in the short term. Irradiation with, for example, UV photons has been employed to produce polar groups at polymer surfaces for the promotion of cell adhesion and spreading.43 UV irradiation in an Ar atmosphere has no significant effect on the adhesion of EC but increases the adsorption of fibronectin.44 Changing the atmosphere to NH3 significantly enhances EC adhesion and spreading.43 Polyurethane in its unmodified state is non-charged but has been modified to become cationic, anionic or zwitterionic.45 These synthetic scaffolds can also be treated with biological molecules such as heparin, growth factors, anti-coagulant peptides and dextran derivatives, and antibiotics. Fibrin or plasma coated ePTFE encourages formation of a virtually complete EC monolayer. Pre-coating PET (polyester) disks with fibronectin, gelatin, collagen I, collagen IV or laminin enhances MSC attachment and influences SMC-like characteristics (Fig. 23.2).

23.3.3 Modified synthetic scaffold materials Biomaterials that can be successfully integrated into vascular grafts should match not only arterial mechanical properties, but also the topography. The cellular response to a biomaterial can be enhanced in synthetic polymer formulations by mimicking the surface roughness created by the ECM components of natural tissue. Synthetic, nano-structured, polymeric biomaterials are being developed that promote cell adhesion and growth for vascular applications. In

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23.2 Mesenchymal stem cells (MSC) grown on PET disks coated with fibronectin or gelatin, in the presence or absence of vascular growth factors. Immunofluorescence analysis of MSC that had been cultured for seven days on PET disks coated with fibronectin or gelatin as indicated. MSC were stained for SM -actin (an intracellular cytoskeletal element) or fibrillin-1 (an extracellular elastic fibre component). Similarities and differences were apparent on these different ECM substrates. MSC cultured on both surfaces exhibited an ordered SM -actin cytoskeleton. Treatment with bFGF resulted in loss of SM -actin staining. Treatment with PDGF-B showed a similar SM -actin cytoskeletal organisation as the `no growth factor' controls. Treatment with TGF 1 resulted in enhanced well-ordered SM -actin staining, especially when grown on fibronectin. MSC cultured on gelatin laid down an extracellular fibrillin network in the absence or presence of growth factors (especially TGF 1 and bFGF). However, MSC cultured on fibronectin failed to deposit wellordered extracellular fibrillin-1.

one such study, poly(lactic-glycolic acid) (PLGA) (50/50wt% mix) was synthesised to possess a range (from micron to nanometer) of surface features. Reduction of surface features was accomplished by treating conventional PLGA with various concentrations of NaOH for select periods of time. Results indicated that NaOH treated PLGA enhanced vascular SMC adhesion and

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proliferation, but decreased EC adhesion and proliferation. Thus, nanostructuring of surface features can significantly improve vascular cell densities. Synthetic materials do not bear epitopes that directly bind cell adhesion receptors, so many studies have focused on modifying polymers using biological or adhesion molecules to ensure cell and tissue biocompatibility. Since integrins are major vascular cell-matrix receptors, attention has focused on attachment of the RGD tripeptide cell adhesion motif to scaffold materials. Adhesion peptides can be stably adsorbed by grafting to the termini of surface-active polymers that adsorb to hydrophobic surfaces46 or to anionic surfaces;47 the surface can then be passivated against non-specific interactions. As well as achieving good surface density of peptides, clustering of peptides is advantageous for cell adhesion.48 An aqueous-based process has recently been reported for surface functionalisation of RGD moieties onto the luminal surface of a prefabricated cardiovascular graft made of poly(carbonate-urea)urethane.49,50 In addition to RGDS, other immobilised cell adhesion peptides (KQAGDV from fibrinogen, recognised by integrin IIb 3; VAPG from elastin) have been adsorbed onto scaffolds. In general, cells adhere more strongly to surfaces modified with these adhesive ligands than to control surfaces.51 Important issues are how adsoption influences short RGD peptide (linear or cyclic) conformation and the specificity of integrin recognition (see section 23.4).

23.3.4 Matrix templates Acellular tubes of extracellular matrix materials based on collagen may be colonised by host vascular cells once placed into a bypassed artery. SMC can also be cultured in vitro for several weeks on tube-shaped biodegradable polymer matrices exposed to pulsatile flow, then seeding the lumen with EC prior to grafting. In general, burst strengths of such grafts are low. Tissueengineered vessels based on rolled up sheets of SMC and fibroblasts embedded within their own matrix to form tubes of media coated in adventitia, have higher burst strength but take at least three months to prepare. Artificial vessels generated by formation of a matrix-rich granulation tissue capsule surrounding biocompatible tubing placed within the peritoneal cavity have high burst strengths, can be prepared within two to three weeks, and are not rejected after transplantation as they are grown in the recipient's own body.52 Several groups are generating vascular grafts containing autologous cells harvested from the vascular system and seeded onto synthetic polymers or bioreabsorbable scaffolds. In our laboratory, we are developing such a vascular graft based on a porous polyurethane elastomeric scaffold pre-seeded with SMC or MSC within biodegradable ECM-based gels and endothelialised prior to preclinical testing. Other groups are using modified polyurethanes such as poly(carbonate-urea)urethane as vascular scaffolds.49 Polyurethanes are elastomeric biomaterials that can be non-charged, or treated to be cationic, anionic, and

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zwitterionic.45 Matsuda et al. recently described a vascular access graft which is a three-layered polyurethane reinforced by spiral polyester fibres.53

23.3.5 Bio-patterning Bio-patterning techniques have been developed that offer the possibility of controlling the chemical surface properties of different types of substrates. For example, ECM protein islands of different geometries and dimensions specific for adhesion of selected cell types, that are able to control cell growth and behaviour have been produced. In this way, self-assembled networks of the basement membrane molecule laminin-1 have recently been obtained on glass substrates by physisorption-assisted microcontact printing.54 Antigenicity was retained as judged by immunofluorescence assays, and the supramolecular organisation of the protein by atomic force microscopy (AFM) revealed the characteristic self-assembling polygonal networks with the well-characterised sub-100 nm quaternary structures of laminin. The formation of these physiological mesh-like matrices involved a one-step soft lithography without preliminary functionalisation of the glass. Recent advances in AFM technology have led to the development of dip-pen nanolithography (DPN) which employs an AFM tip as a `nib', a solid substrate as the `paper' and molecules with a chemical affinity for the substrate as the `ink'.55 DPN is able to pattern substrates with linewidths in the range 10±100 nm, generate biologically active multicomponent nanostructures and produce arrays through the use of multiple probes.56,57 Thus, it should be possible to exploit these technological advances to introduce ordered ECM signals into grafts and other biomolecular devices.

23.3.6 Vascular cells used to seed vascular grafts EC and SMC sources for replacement tissue engineered grafts include (autologous or heterologous) omentum microvasculature, saphenous veins, and (heterologous) umbilical vein or artery. In general, cells from elderly patients frequently do not grow well or deposit significant functional vascular ECM. Several groups have explored the use of adult stem cells in vascular tissue engineering. In our laboratory we are using commercially available coronary artery and aortic SMC, and MSC from bone marrow aspirates to seed graft scaffolds. We are also using porcine SMC and MSC in pre-clinical testing of graft `take' and patency.

23.4 Coating ECM molecules on surfaces ± a cautionary tale Understanding the features of the individual cells of blood vessels and their surrounding matrix is critical to biomimetic design in vascular tissue engineering. It is also essential to know the chemical and physical characteristics

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of different synthetic surfaces, and to ensure that an appropriate conformation is achieved when attaching ECM molecules to different surfaces. Several recent studies provide critical insights into how adsorption of proteins to different surfaces can affect their conformation and biological functions such as cell adhesion.

23.4.1 Adsorbing proteins onto surfaces The exposure of synthetic materials to solutions rich in biological molecules often triggers a dynamic non-specific adsorption process (physisorption) in which surface/solution chemistries and molecular structural rearrangements play a role.58 Physisorption of proteins to substrates is driven by the electrostatic double-layer force (EDL force), the hydrophobic effect and relatively weak van der Waals forces.59,60 The EDL force is dependent on the concentration and valency of charged solutes and the surface charge density of both substrate and biomolecule.60 Solute concentration and valency have been shown to play an important role in the adsorption of negatively charged molecules such as DNA to negatively charged substrates such as mica.61 The adsorption of biomolecules such as proteins to hydrophobic surfaces may involve van der Waals forces, hydrogen bonds and the formation of a water-free contact layer allowing the protein to unfold without solvating all of its hydrophobic residues.58,62 Adsorption mediated molecular structural rearrangements, influenced by surface chemistry, have been demonstrated for synthetic and biological macromolecules. On hydrophilic mica substrates adsorbed synthetic ethylenediamine core poly(amidoamine) dendrimers appeared as flattened disc-like structures but on hydrophobic substrates dendrimer height approximated theoretical diameter predictions.62 Similar substrate dependent morphological changes have been observed for diverse biomolecules such as Alzheimer's -amyloid peptides and von Willebrand factor (vWF).63,64 Substrate dependent morphological and functional changes have also been observed for important constituents of the ECM. Fibronectin is a major adhesive glycoprotein, which binds to many ECM components and cells via membranebound integrins. Adsorption of fibronectin to different substrates is thought to modulate cell adhesion, spreading and migration through changes in protein conformation.65,66 AFM studies of fibronectin identified an extended morphology on hydrophilic and a condensed morphology on hydrophobic substrates.67 An adsorption study of fibronectin to self-assembled monolayers of alkanethiols with defined surface chemistries determined that although functional activity was dependent on structural changes, there was no simple correlation with substrate hydrophobicity.68 Substrate-dependent alterations in fibronectin mediated cell binding and spreading may be due to the relative movements of the RGD loop on Fn-III9 and the synergy site on Fn-III10. Steered molecular dynamics simulations predict a relative movement of these two

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domains under mechanical stress of 2.3 nm, which would preclude integrin binding at both sites.69 Substrate-induced conformational changes may generate similar relative domain movements. Elastic fibres are a composite biomaterial composed of elastin and fibrillin microfibrils, which are themselves heteropolymers.7,70 Within the elastin sequence, distinct exons encode for alternating hydrophobic and cross-linking domains. Small elastin peptides derived from exons 20, 21, 23 and 24, coacervate in heated solutions to form fibrillar aggregates71±73 (see section 23.5). In a variable temperature in situ AFM study these elastin peptides selfassembled in a substrate-dependent manner. On hydrophilic mica surfaces the peptides adsorbed as discrete, rounded aggregates. Adsorption to hydrophobic highly ordered pyrolitic graphite (HOPG) induced a fibrillar arrangement.74 The authors suggest that the order observed on HOPG substrates results from hydrophobic peptide-substrate interactions, which form an energetically closepacked arrangement acting as a template for fibril growth. Recent AFM investigations in our laboratory have highlighted the roughness characteristics of several materials (Fig. 23.3), and demonstrated that fibrillin microfibrils also exhibit substrate-dependent morphologies.75 Fibrillin microfibrils adsorbed to a relatively hydrophobic substrate (mica coated with poly-L-lysine) had a compact morphology and a directional asymmetry to the bead structure. On adsorption to hydrophilic mica, however, the microfibrils appeared more diffuse and the bead asymmetry was lost (Fig. 23.4). Biochemically distinct type VI collagen microfibrils are co-purified with fibrillin microfibrils from most tissues following bacterial collagenase digestion and size exclusion chromatography. These supra-molecular assemblies form an extensive microfibrillar network linking cells and many ECM components and maintaining the integrity of tissues such as blood vessels, lung and skin.3 The classical double-beaded appearance of type VI collagen microfibrils was evident on a hydrophobic substrate (mica coated with poly-L-lysine). Adsorption to hydrophilic mica severely disrupted microfibril morphology inducing a major conformational re-organisation along the whole collagen microfibril repeat (Fig. 23.4). Plasma fibronectin is thought to have a compact conformation; the extended fibronectin morphology observed on hydrophilic surfaces may be due to disruption of inter-domain ionic interactions. The relatively weak van der Waals forces and hydrogen bonds which are thought to dominate sample-substrate interactions on hydrophobic substrates,62 may not interfere with intermolecular ionic interactions reducing substrate-induced distortion of the sample. Therefore fibronectin and ethylenediamine core poly(amidoamine) dendrimers may adsorb to hydrophobic surfaces with a more `native' structure.62,67 Preservation of biological function is a vital consideration in the coating of synthetic materials with biomolecules, and the interaction of cells with synthetic polymers with physico-chemical surface properties that strongly influence cell

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23.3 Roughness of non-biological surfaces visualised by tapping mode AFM. Tapping mode AFM scans (5 m  5 m) and height profiles. (a) Mica minerals are layered crystals in which tetrahedral sheets of (Si, Al)2O5 are ionically linked by central Al2(OH)2 layers. The hydrophilicity and atomic scale flatness (RMS roughness ˆ 0.22 nm) of mica make it a suitable substrate for the adsorption of bio-molecules prior to imaging by AFM and EM techniques. (b) Tissue culture plastic has a hydrophilic surface which enhances cell adhesion and a higher RMS roughness value than mica (2.54 nm) which limits the ability of the AFM to image adsorbed bio-molecules. (c) Nitrocellulose membranes used for blotting proteins and nucleic acids are extremely rough compared to mica and tissue culture plastic (RMS roughness ˆ 80.65).

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23.4 Substrate-dependent morphologies of fibrillin and type VI collagen microfibrils. Tapping mode AFM micrographs (2 m  2 m) and microfibril height contour maps (1 m  0.16 m) of microfibrils adsorbed to hydrophilic (mica) and hydrophobic (mica ‡ poly-L-lysine) substrates. (a) Fibrillin microfibrils adsorbed to poly-L-lysine coated mica were closely packed with a pronounced bead shoulder region evident in the height contour map (arrows). On mica, the shoulder region was absent and the microfibrils appeared more diffuse. (b) The characteristic double bead (braces) of type VI collagen microfibrils was preserved on mica coated with poly-L-lysine but on the more hydrophilic mica double beaded-regions were rarely identifiable.

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morphology and growth has many potential applications in medicine and biotechnology (see sections 23.3.2 and 23.3.3). Thus, in the light of recent studies, it is essential to consider the surface chemistry of synthetic materials in relation to the biological function of adsorbed molecules.

23.5 Biological seeding materials The ECM is the authentic substrate for connective tissue cells including SMC and EC. It is a three-dimensional multimolecular milieu with complex higherorder architecture and tissue-specific compliant properties. Cells cultured within such a biological matrix experience a richer, more complex physical environment and markedly different geometry than cells on flat surfaces. Cells sense their surrounding ECM, are directly anchored to it, respond to ECM through cell-matrix adhesion signals that influence cell growth, migration, differentiation, survival, tissue organisation, and remodel their surrounding matrix. Mechanical signals feed back to modulate cell phenotype in an interative process. Three-dimensional vascular ECM materials are thus powerful tools for modulating grafts and regulating vascular cell phenotype within them.76±78 A number of ECM-based cell seeding biomaterials have been developed that can be cast as gels or membranes and used in the biological engineering of vascular and other connective tissues. Below are summarised several key materials. Collagen gels Collagen is the most abundant protein in the body, provides tensile strength to tissues, and has long been an attractive biomaterial for use in tissue engineering. Three-dimensional collagen gels based on rat tail tendon collagen or foetal bovine skin collagen have been extensively used in cell biology over the past three decades.39,79 Animal or human tissue-purified procollagen solutions are reconstituted into gels by changing the pH and temperature, but there are concerns about immunogenicity and disease transmission. Future use of recombinant fibrillar collagens I and III should circumvent this issue. Collagen gels have been used to study mechanisms of cell migration and the influence of the three-dimensional state on cell behaviour.78 The physical properties of collagen gels can be profoundly influenced by crosslinking using glutaraldehyde or specific chemical crosslinkers.80,81 Chemical glycation procedures have been developed to adjust, in a controllable manner, the elastic character of collagen gels. Other crosslinking strategies include use of transglutaminase which catalyses formation of an isopeptide -glutamyl--lysine bond between or within polypeptide chains,82 or lysyl oxidase which is a physiological copper-dependent amine oxidase.83 Collagen gels can be cast as liquid crystalline collagen arrays that modify cell migration.84 Electrospinning of collagen nanostructures has also been

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described.85 When not constrained at their edges, collagen gels contract under the influence of cell-generated forces, thereby altering their mechanical properties.86 In vascular tissue engineering, collagen gels were used to seed SMC onto scaffolds such as Dacron.87 However, the cells did not proliferate significantly and failed to lay down vascular ECM. In order to mimic the vascular medial elastic fibre-rich ECM environment, we have modified collagen gel contraction with recombinant RGD and non-RGD-containing fibrillin-1 fragments and with fibronectin. We also found that elastic fibres are deposited at the periphery of these collagen gels, where the SMC are arranged circumferentially and are under tension (Fig. 23.5). Future ECM modifications could include incorporation of ECM polymers such as fibrillin or collagen VI microfibrils that can modify the gel mechanical properties34 and provide linear repeating RGD motifs for SMC attachment and alignment.3,28 Collagen-GAG composite gels have also been developed.88,89 A construct-sleeve hybrid graft has been described which uses a biological support sleeve, based on collagen I gels crosslinked with glutaraldehyde or ultraviolet, to provide temporary reinforcement during cellmediated construct remodelling.90 Fibrin gels Fibrin gels exploit the final stage of the coagulation cascade in which fibrinogen molecules are cleaved by thrombin and form three-dimensional fibrin gels. Fibrin forms a complex network with fibril structures and crosslinking characteristics determined by the polymerisation conditions. Human fibrin is clinically available from autologous sources and cryoprecipitated pooled blood plasma. Fibril gels can be injected into porous scaffolds, or other constructs, with SMC or MSC (inject fibrinogen plus thrombin). They have advantages of biocompatibility, biodegradability and haemostasis, and MSC and fibroblasts can proliferate well within them and migrate into neighbouring tissues.91 Factor XIII (a transglutaminase) can be used to crosslink these gels.92 Fibrin vascular `media equivalents' have been developed that form a compact strong structure having, after six weeks, a tensile modulus and ultimate tensile strength similar to that of rat abdominal aorta.93 Elastic fibre deposition has been observed in cultures of SMC within fibrin gels.94 Hyaluronan supports Hyaff-11Õ (Fidia Advanced Biosystems) is a benzyl esterified form of hyaluronan (HA)95 that forms a hydrophobic polymer than can be spun or woven to form a scaffold for cell growth. Hyaff-11Õ has a degradation time of ~40 days, and does not induce an inflammatory response or fragmentation of the scaffold. Instead, the material becomes more hydrophilic, forming a gel similar

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23.5 SMC contraction of collagen gels and deposition of elastic fibres. (a) Contraction of rat tail type I collagen gels over time, containing 0, 10,000, 20,000, 40,000 or 80,000 human coronary artery SMC per gel. (b) Phase contrast microscopy of human coronary artery SMC at day 21 at the centre and periphery of collagen gels. Arrows indicate orientated cells around the circumference of the gel. (c) Elastin (Millers) staining of collagen gels with 20,000, 40,000 or 80,000 cells per gel, showing elastin deposition at the periphery of the gel.

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23.6 Environmental scanning electron microscopy (ESEM) of MSC cultured on Hyaff-11Õ ESEM was used to investigate the deposition of ECM by MSC cultured on Hyaff-11Õ. By gradually reducing the pressure within the ESEM microscope chamber, ECM fibrils could be visualised in a native hydrated state (a±c). Bars = 50 m.

to native hyaluronan found in the ECM. A recent study has investigated the use of two forms of Hyaff-11Õ, unpressed and pressed fibrous non-woven Hyaff11Õ scaffolds, as scaffolds for EC in tissue engineered vascular grafts.96 EC attached to the individual fibres of Hyaff-11Õ, and deposited an organised subendothelial matrix containing laminin, fibronectin, collagen IV and collagen IV. Thus, Hyaff-11Õ based biopolymers may provide a suitable surface to promote endothelialisation within vascular grafts. We have also shown that vascular SMC and MSC adhere well to Hyaff-11Õ, so Hyaff-11Õ may also be suitable for incorporating into vascular medial layer scaffolds (Fig. 23.6). Elastin-based materials Elastin is a polymeric structural protein with alternating hydrophobic and crosslinking domains that imparts the physical properties of extensibility and elastic recoil on arteries.7 It has an intrinsic capacity to form an ordered assembly through a process of hydrophobic self-aggregation or `coascervation', in which the protein comes out of solution as a second phase on an increase in solution temperature. The kinetics of the transition appear to be that of a nucleation process. The temperature at which this transition takes place is dependent on elastin concentration, ionic strength and pH. Recently, the coascervation behaviour of recombinant fragments comprising as few as five domains of human tropoelastin shows alignment of lysine residues for crosslink formation and self-assembly into crosslinked matrices with solubility and mechanical properties similar to native elastin.71±73 Elastin-based materials can be cast as membranes and tubes, and are suitable for use in composite vascular grafts. We have produced aligned arrays of isolated fibrillin microfibrils and are developing novel microfibril-elastin composite materials in vitro (Fig. 23.7).

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23.7 Fibrillin microfibrils can form networks in vitro, as observed by STEM. Composite image of 1.3 m  1.3 m STEM scans (a) and extracted 1 m  1 m regions (b and c). The organisation of fibrillin microfibrils into ordered sheets has been observed at low NaCl concentrations.97 (a) and (b) Treatment with a nonionic detergent (Triton X-100) induces similar ordered sheet-like structures aligned laterally at the microfibril beads (arrows). (c) Lateral association of fibrillin microfibrils has also been observed following incubation with recombinant elastin peptides. Sheet-like structures of recombinant elastin peptides (arrow) link two fibrillin microfibrils.

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Hydrogels Hydrogels are crosslinked hydrophilic polymer networks that may absorb more than 1000 times their dry weight in water, giving them physical characteristics similar to soft connective tissues.78,98±99 These nominally biocompatible materials are ideal for many clinical applications. Swelling and hydration occur without dissolution of the polymer since the process of crosslinking forms an insoluble network. Hydrogels are highly permeable, which facilitates exchange of oxygen, nutrients, and other water-soluble metabolites. They offer potential for sustaining viable vascular cells within graft scaffolds and production of three-dimensional vascular matrices.

23.5.1 Poly(ethylene glycol) (PEG)-based hydrogels Biomaterials based on PEG are normally highly resistant to protein adsorption and cell adhesion, so cell-material interactions are limited to the biomolecules that are covalently incorporated into the material.78,100 Protocols have been devised for the covalent conjugation of RGD to PEG hydrogels,46 and incorporation of vascular growth factors into PEG hydrogels.101 The elastinderived peptide VAPG, which is a useful vascular cell adhesion ligand, has been covalently immobilised via a flexible linker chain into photopolymerisable hydrogels based on acrylate-terminated derivatives of PEG. One study showed that VAPG was specific for SMC adhesion and spreading in a ligand concentration dependent manner; fibroblasts, EC and platelets could not adhere to it.76 In situ crosslinking of a biomimetic peptide-PEG hydrogel via thermally triggered activation of factor XIII has also been described,102 as well as artificial ECM protein hydrogel films prepared by isocyanate crosslinking.103

23.5.2 Crosslinkable hyaluronan hydrogels HA is a major constituent of the developing vascular ECM, and is the only nonsulphated glycosaminoglycan (GAG). It is synthesised at the cell membrane directly into the EC space, is polyanionic and highly hygroscopic, biocompatible and biodegradable, and performs important biological functions such as stabilising and organising the ECM, regulating cell adhesion and motility and mediating cell proliferation and differentiation. Osmotic swelling of HA provides compressive strength. New crosslinking strategies have been developed to prepare HA-based hydrogels, including disulphide crosslinking of thiolated HA derivatives, generated by addition of thiols to , -unsaturated esters and amides of PEG, that are cytocompatible with potential for injectable cell delivery.104 Another approach is photopolymerisable HA-based hydrogels copolymerised with poly (ethylene glycol) diacrylate (PEG-DA), into which can be incorporated RGD or other biological motifs.105 Hyaluronan micro-

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spheres conjugated to specific cell surface antigens show selective binding to cells expressing these receptors, and are suitable for site-specific targeting.106

23.5.3 Alginate hydrogels Alginic acid, a polysaccharide from seaweed, is a family of natural copolymers of -D-mannuronic acid and -L-guluronic acid. Aqueous solutions of the sodium salt of alginate gel upon contact with calcium ion to form a stable elastic gel. Because of their biocompatility, abundant source and low cost, they have been widely exploited in tissue engineering. Alginates crosslinked with calcium sulphate have been used as cell delivery vehicles, although cell viability during the gelation process is difficult to control.107,108 Modified alginate gels may have potential uses in vascular tissue engineering.

23.6 ECM-regulated delivery of therapeutic growth factors In natural vessels, growth factors influence vascular ECM synthesis and deposition and are, in turn, regulated by ECM. Cross-talk between growth factor and integrin cell-matrix receptors, and their signalling pathways, also directly and indirectly modifies growth factor actions. Thus it is important that strategies are devised that allow the controlled release of vascular growth factors during graft engineering. Bioactive polymers are now being engineered that not only provide physical support and a template for cells, but also closely mimic the in vivo release mechanisms of vascular growth factors from the ECM.109 Biomaterials allow sequential delivery of growth factors that have different specialised roles in blood vessel formation and function. Key growth factors in vascular tissue engineering include TGF 1 which stimulates SMC proliferation and ECM deposition, PDGF-BB which helps to stabilise newly formed capillaries by recruiting SMC and regulating their contractile phenotype, bFGF which influences PDGF receptor levels and EC and SMC behaviour, and VEGF which initiates endothelial capillary formation. Some ECM-based growth factor release systems are outlined below. Bioactive VEGF-modified hydrogels have been developed that allow growth factor release, upon local cellular demand, as cells remodel the hydrogels.109 bFGF and VEGF can be bound to heparin or fibronectin for slow release in situ. pH can be used to regulate VEGF binding to fibronectin and fibronectin-heparin matrices. VEGF binding is increased at acidic pH, and these interactions are enhanced by heparin; VEGF is then released as the pH is raised, and retains its ability to stimulate EC.110 Proteolysis of membrane-anchored heparin-binding EGF-like growth factor via metalloprotease (MMPs, ADAMS) activation yields HB-EGF which plays a key role in wound healing;111 these release principles are now being developed for vascular constructs. Another ECM polymer with strong

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potential for use in controlled growth factor release is the fibrillin microfibril, and associated with latent TGF binding proteins (LTBPs) which bind TGF within the cell, regulates its secretion, extracellular deposition and activation.23,35 Microspheres loaded with bFGF can be delivered into the circulation for uptake by the surrounding EC.109 Microparticles of heparin and alginate crosslinked gel can be used as injectable artificial matrices to stabilise bFGF and induce angiogenesis by controlling its release.112 Controlled TGF 1 release from biodegradable polymer microparticles based on PEG/PLGA113 or polycaprolactone that are implanted into regenerating tissue or tissue engineered constructs, has also been described. IGF-1 and TGF 1 can be released in a controlled manner from PLGA microspheres within a photopolymerising hydrogel; this method allows delivery of growth factors within porous hydrogels.114 The release kinetics of TGF 1 from fibrin clots using liposome encapsulation show this to be another promising slow-delivery system for tissue engineering.115

23.7 Future trends Mimicking the vascular ECM is certain to be a major theme of future vascular tissue engineering strategies. Current developments include use of correctly folded recombinant ECM molecules and domains that present RGD and other cell adhesion motifs in biological context, thereby ensuring specific vascular cell signalling responses. Bio-patterning and imprinting of ECM molecules onto scaffold surfaces, or within composite layered scaffolds, provide opportunities to direct vascular cell migration into graft scaffolds and then regulate their phenotype. The development of scaffold materials that allow adsorption of ECM molecules and polymers without loss of biological function, and the use of tropoelastin to coat surfaces to encourage EC attachment but inhibit platelet adhesion, will be valuable advances. Novel recombinant ECM molecule based seeding materials are being developed that can be assembled in vitro according to the natural principles of ECM assembly, e.g., controlled enzymatic processing of collagens and fibrillin. There is a need for chemical crosslinking strategies that do not alter biological signals.

23.8 Acknowledgements Research from our laboratory described in this chapter was funded by the UK Centre for Tissue Engineering (MRC, BBSRC, EPSRC), and an MRC programme grant (CMK). We thank Dr D. Lee who conducted the immunofluorescence analysis of mesenchymal stem cells, Dr S. Sinha who generated the micrograph of the porcine coronary artery, and Dr Carolyn J. P. Jones who expertly conducted the transmission electron microscopy.

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24

Biomineralisation processes F C M E L D R U M , University of Bristol, UK

24.1 Introduction Biomineralisation describes the production of inorganic solids by organisms, and is extremely common in nature, with over 60 biominerals having been identified among all five of the animal kingdoms (Lowenstam, 1989; Mann, 2001). Of known minerals, approximately 20% are amorphous and 80% crystalline, although the number of amorphous minerals may be an underestimate due to the problems of detecting amorphous materials in the presence of crystalline ones (Weiner, 1997; Addadi, 2003). Biominerals are typically characterised by unique and elaborate morphologies, and display properties optimised for their function. For example, biominerals fulfilling structural roles typically possess remarkable mechanical properties which can rival those of engineering materials. The magnetite crystals providing magnetotactic bacteria with a magnetic dipole and enabling navigation in the earth's magnetic field are always a single magnetic domain in size. Of note also is the ability of organisms to selectively precipitate minerals such as SrSO4 and BaSO4 that are significantly undersaturated in the surrounding environment of the organism (Mann, 2001). This chapter will review the principal mechanisms used by organisms to control the formation of crystalline and amorphous minerals, examining in particular control of phase, orientation, morphology and mechanical properties. Many of the mechanisms discussed will be based on the study of calcium carbonate. Calcium-based minerals represent approximately half of known biominerals and include the most abundant species calcium carbonate and calcium phosphate, which are important in invertebrate and vertebrate skeletonformation respectively. Iron-based biominerals are also significant, ranging from the ferrihydrate core of the ferritin protein (Chasteen, 1999) to the magnetite crystals which create the magnetic dipole within magnetotactic bacteria and will be discussed (Bazylinski, 2003, 2004). Of the amorphous biominerals, amorphous silica is by far the most significant, being widespread among plants, Protoscista such as diatoms and radiolaria, and occurring in some lower animals

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such as sponges. The diatoms, for example, are extremely abundant in freshwater and marine environments, and are estimated to be responsible for 20-25% of all organic fixation is carried (BruÈmmer, 2003) and the mechanisms involved in the precipitation of their silica skeletons have been investigated in some detail (KroÈger, 2001, 2002). Precipitation of silica by organisms will be contrasted with that of amorphous calcium carbonate, which is characteristically short-lived in synthetic systems.

24.2 `Biologically-induced' and `organic matrixmediated' mineralisation Biomineralisation has been divided into two fundamentally different classes, based on the degree of biological control (Mann, 2001). The precipitation of a mineral as a result of interactions between an organism and the environment, when little control is exerted over the mineralisation process is termed `biologically-induced' mineralisation. An example of this type of mineralisation is the decoration of bacteria by a wide range of minerals including iron and manganese oxides. The surface of the bacteria contains a range of macromolecules which can express reactive groups such as carboxylates and phosphates capable of binding metal ions present in the environment. The bound metal ions then serve as a nucleation site for precipitation, where local supersaturation is induced by changes in the pH or redox potential at the surface due to the bacterial metabolism, or by the production of reactive ligands such as sulphide (BaÈuerlein, 2003). In `organic matrix-mediated' mineralisation the organism directly controls mineralisation, having evolved specific strategies to produce minerals of specific size, morphology, structure and orientation. This process is of considerably more interest as the techniques used by organisms can provide inspiration for synthetic crystal growth experiments, and will form the basis for this review. As organisms clearly cannot manipulate parameters such as temperature or pressure, as is commonly done synthetically to control crystal growth, the strategies they use rely on organic molecules to control mineralisation; confining a space, forming an organic matrix framework, controlling ion input, constructing a nucleation site, controlling crystal orientation and growth and terminating crystal growth. Some or all of these may be involved in the precipitation of a given mineral (Lowenstam, 1989; Mann, 2001).

24.3 Organic macromolecules The organic macromolecules involved in regulating biomineralisation processes can be categorised either as insoluble matrix molecules or soluble `control' macromolecules (Weiner, 1997). Biologically controlled crystal growth always takes place within a designated space which is delineated by a structured organic

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matrix. This can vary from simple intracellular lipid vesicles in the case of magnetosomes in magnetotactic bacteria, to more complex extracellular macromolecular frameworks such as the organised collagen fibrils in bone. This environment is crucial to the control of mineralisation by organisms. It determines the location of mineral deposition and enables the organism to control the composition of the solution in terms of the supersaturation of the desired mineral and the presence of any additional soluble additives. The threedimensional form of the mineralisation site, which often changes during the mineralisation process can also dictate the final morphology of the mineral structure formed due to spatial constraint (Addadi, 2003; Park, 2004; Mann, 2001). The organic framework is often further functionalised by adsorption of the soluble control macromolecules to provide an additional mechanism for controlling crystal nucleation and growth (Hunter, 1996; Weiner, 1997). There are many examples in biology of orientated crystal growth, and there is some evidence that this may be induced by organisation of the nucleating organic matrix. The rate of nucleation of calcium carbonate can also be increased, and the morphology altered, by immobilised control macromolecules, while the same proteins acting from solution can inhibit crystal growth (Wheeler, 1981, 1984, 1991). The difference in behaviour between the immobilised and free protein can be attributed to conformational changes and rigidity on binding to a solid substrate. In addition to controlling crystal growth when bound to a solid substrate, control macromolecules may also be employed as soluble additives, adsorbing to a crystal during growth and modulating morphologies and textures (Albeck, 1993, 1996a,b; Aizenberg, 1994). This is supported by the observation that biominerals can contain organic macromolecules within the mineral, and that precipitation of calcium carbonate in the presence of macromolecules extracted from selected calcium carbonate biominerals results in interaction with specific crystal planes. Intracrystalline macromolecules have perhaps been best studied in biogenic calcium carbonates and calcium phosphates, the majority of which contain organic molecules at levels of up to a few wt%. Demonstration of the presence of intra-mineral macromolecules and their subsequent characterisation is dependent on the ability to dissolve the mineral without damaging the macromolecules and then isolate the macromolecules. This is readily achieved for Ca-based minerals which can be dissolved most successfully by incubation with ion-exchange resin (Albeck, 1996b; Gotliv, 2003), but is more challenging for silicas which are typically dissolved using HF. Despite this relatively extreme treatment, organic macromolecules have been isolated from a range of silica biominerals, including diatoms (KroÈger, 2000, 2001), plants (Perry, 2003) and sponge spicules (Cha, 1999; Shimizu, 1998). Extraction with an aqueous solution of ammonia fluoride has been demonstrated to provide a more gentle method for dissolution of diatom biosilica (KroÈger, 2002).

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The control macromolecules present in calcium carbonates are usually highly acidic (Lowenstam, 1989). The total assembly of macromolecules isolated from calcite and aragonite biominerals are rich in aspartic acid and glutamic acid, and frequently contain bound polysaccharides (Albeck, 1996a; Weiner, 1997). Purification and characterisation of these macromolecules has proven extremely challenging, due in part to their high charge, although some common partial sequences of amino acids have been identified, consisting of runs of poly-Asp and alternating sequences with poly-Asp at every other residue, commonly separated by either glycine or serine (Weiner, 1975). This sequence exhibits regular repeating negative charges which may bind Ca2+ ions and could be significant in controlling crystal growth. In addition, poly(aspartic acid) can adopt a -sheet structure on binding to Ca, which appears to be important in enabling structured interaction with crystal faces (Addadi, 1987; Weiner, 1975). Application of a novel gel electrophoresis fixing and staining protocol has recently offered a route to visualisation and separation of all components of the macromolecule assembly (Gotliv, 2003). Indeed, only one fraction of the macromolecules extracted from the aragonite layer of the mollusk Atrina rigida was active in inducing precipitation of aragonite within a chitin framework. Dissolution of the silica framework produced by diatoms with HF yields a species-specific set of low molecular mass proteins termed silaffins, together with large quantities of long-chain polyamides (KroÈger, 1999, 2000, 2001). The silaffins-1A and -1B isolated from Cylindrotheca fusiformis are polycationic and contain repeated pairs of lysine residues which are converted into three derivatives, -N-dimethylysine, phosphorylated -N-trimethyl--hydroxylysine and lysines covalently linked to long-chain polyamines comprising 6 to 11 units of N-methyl-propylamine (KroÈger, 2001, 2002). More gentle dissolution procedures using an aqueous solution of ammonium fluoride enabled the silaffin proteins to be extracted in their native state, such that the serine groups are phosphorylated (KroÈger, 2002). A second silaffin protein, termed silaffin-2 has also been extracted from C. fusiformis. Silaffin-2 is in contrast polyanionic in structure and also bears unusual amino acid modifications (Poulsen, 2003). NatSil-1A and the long-chain polyamines are extremely active in promoting silica precipitation in vitro and are anticipated to have a similar role in vivo (KroÈger, 2002; Sumper, 2003). NatSil-2 alone does not precipitate silica from a silicic acid solution in vitro, but is active in combination with long-chain polyamines, rapidly precipitating silica under conditions where neither of these organic components would do so alone. Beyond a critical concentration of natSil-2, further increase in the natSil-2 concentration inhibited precipitation. The silica precipitation activity of natSil-1A was also impeded by natSil-2, probably due to electrostatic shielding of the polycationic amine chains. NatSil2 may therefore function as a regulator of the silica precipitation behaviour of long-chain polyamines and natSil-1A and may be active in silica morphogenesis. While large interconnected spherical or pear-shaped silica particles

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24.1 Silica precipitates formed by mixtures of natSil-1A and natSil-2 (a) natSil-1A at 0.5 units/l, natSil-2 at 0.3 mM; (b) natSil-1A at 5.0 units/l, natSil-2 at 0.3 mM; (c) natSil-1A at 2.0 units/l, natSil-2 at 0.3 mM; (d) natSil-1A at 1.6 units/l, natSil-2 at 0.2 mM. Scale bars ˆ 2 m. Reproduced with permission from Poulsen, 2003, Copyright 2003, National Academy of Sciences, USA.

were produced in the presence of either low or high natSil-1A/natSil-2 ratios, intermediate ratios yielded porous silica blocks permeated with 0.1±1.0 m pores (Fig. 24.1). These structures may derive from assembly of the organic phase in the presence of the polysilicic acid molecules such that it provides a template for the ultimate form of the mineral phase (Poulsen, 2003). Silica sponge spicules occlude a central protein filament, termed a silicatein, which can again be isolated by dissolution of the spicules in HF (Shimizu, 1998; Cha, 1999). The silicatein protein filaments are in the order of 1±2 mm long and 30 m in diameter and exhibit a regular repeating structure comprising three similar subunits, termed -, - and -silicatein. Analysis of the amino acid composition of these proteins demonstrated approximately 25% of the hydroxy amino acids serine, threonine and tyrosine which typically appear in clusters, 15% glycine and about 20% acidic amino acids (Shimizu, 1998). The hydroxyl rich structure is postulated to be important in the silicification process. Indeed,

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extracted silicatein proteins are active in directing the polymerisation of silica and silicones in vitro, under conditions where an acid or base catalyst would be required in their absence (Cha, 1999). This conformation, as well as the density of hydroxyl residues is likely to be important in this process.

24.4 Control over crystal structure Given the enormous number of recognised biominerals, and the fact that organisms frequently precipitate minerals from environments that are highly undersaturated with respect to those minerals, it is clear that organisms actively select these minerals. This ability is based on the presence of ion-specific pumps and channels which determine the composition and concentration of ions in the mineralisation space, and can be used to dictate the order that ions are introduced. The reactant cations and anions can also be kept separately until precipitation is to be initiated (Lowenstam, 1989). Many minerals exist in a number of closely related forms, such as the iron oxides and hydroxides or the polymorphs of calcium carbonate. Control over the iron oxides/hydroxides produced can be achieved by determining the composition and pH of the solution from which the crystal precipitates, and through interaction with the surrounding organic matrix. The protein ferritin comprises a spherical protein shell, which encapsulates a core of ferrihydrite, 5Fe2O3.9H2O (Fig. 24.2). Fe(II) is bound at a ferrioxidase site located within an H-chain subunit where oxidation to Fe(III) is catalysed, prior to transport to a nucleation site comprising six negatively charged groups on the inner surface of the protein shell. Hydrolysis of Fe(III) ions at the nucleation site generates the poorly ordered ferrihydrite mineral (Chasteen, 1999; Harrison, 1996). In the case of magnetotactic bacteria, the magnetite crystals which give the bacteria their overall magnetic dipole are formed within separate vesicles termed magnetosomes (Fig. 24.3). The magnetosome membrane contains at least one major protein with molecular weight 22± 24 kDa which appears common to all strains. This specific protein is likely to have specific functions in the accumulation of iron, nucleation of the iron oxides and in redox and pH control (SchuÈler, 1999). Recent experiment has shown that the magnetosome vesicles form prior to magnetite formation, and that once initiated, magnetite biomineralisation occurs simultaneously in many vesicles (Komelli, 2004). Evidence from MoÈssbauer studies (Frankel, 1983) and high resolution transmission electron microscopy (Mann, 1984) has suggested that the magnetite crystals may not precipitate directly, but form via a low-density hydrous Fe(III) oxide and ferrihydrite precursors. It was therefore proposed that Fe(III) is taken up by the cell, where it is reduced to Fe(II) as it enters the cell. Subsequent re-oxidation then yields a low-density hydrous Fe(III) oxide which is dehydrated to form crystalline ferrihydrite. Finally, within the magnetosome, one third of the Fe(III) ions are

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24.2 Transmission electron micrograph of the iron storage protein ferritin negatively stained with uranyl acetate (an individual protein molecule with core is arrowed). The inset shows a schematic diagram of the hollow protein shell and ferrihydrite core.

reduced and dehydration yields magnetite (Frankel, 1983). Recent experimental evidence obtained from analysis of magnetite deposition in M. gryphiswaldense has, however, indicated that in this species, Fe(III) is taken up and rapidly converted into Fe3O4 in the absence of a precursor phase (SchuÈler D, 1998). Calcium carbonate can exist in the anhydrous crystalline polymorphs calcite, aragonite and vaterite, and as there are no examples of transformation between calcite and aragonite after precipitation selection of either calcite or aragonite must occur at nucleation (Weiner, 1997). Calcite and aragonite are close in stabilities, although calcite is the most stable polymorph under standard conditions, while vaterite is unstable and rare in nature (Lippmann, 1973). Selection of polymorph at nucleation could be achieved either by inhibiting the nucleation of a stable form, or via preferred nucleation of a metastable form. The former mechanism is achieved using organic molecules, which bind to crystals of the stable, but not the metastable form. A co-operative mechanism may exist where the substrate proteins are responsible for nucleation and orientation, while an inhibitor of the stable polymorph selects the metastable polymorph (Weiner, 1997). A number of experiments have suggested that soluble macromolecules are involved in the selection of calcite or aragonite (Belcher, 1996; Falini, 1996; Gotliv, 2003; Levi, 1998). The structure of the organic matrix present in the nacreous layer of molluscs was mimicked in vitro using layers of -chitin and silk fibroin and soluble macromolecules extracted either from the calcitic or aragonitic layers of mollusc shells (Falini, 1996; Levi, 1998; Gotliv, 2003). Deposition of calcium carbonate on the complete matrix assembly induced

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24.3 Transmission electron micrograph of a magnetotactic bacterium designated MV-2 (Meldrum, 1993) negatively stained with uranyl acetate showing the chain of magnetite crystals which provides the bacterium with an overall dipole moment (arrowed).

calcite growth inside the chitin when calcitic macromolecules were used, and aragonite when the macromolecules had been extracted from an aragonite layer. The specificity of these macromolecules was achieved only with the complete substrate assembly. The macromolecules which promoted calcite nucleation are strongly polyanionic and more strongly acidic than aragonite-inducing ones. They may have provided a strong binding site for Ca2+ ions, creating a high local supersaturation. In contrast, the macromolecules that promote aragonite nucleation are unique to the aragonite layer. This system has also provided a suitable assay to test the function of purified protein fractions extracted from calcite and aragonite biominerals (Gotliv, 2003). A clear demonstration of the role of soluble macromolecules in controlling crystal structure comes from the stabilisation of amorphous calcium carbonate (ACC) by organisms (Addadi, 2003). ACC is a strongly hydrated, amorphous form of calcium carbonate that is highly unstable towards crystallisation.

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Extraction of the glycoproteins occluded within ACC in sponge and ascidian spicules showed these to be rich in glutamic acid and/or glutamine and quite distinct from those within the calcite component of the Clathrina spicule (Aizenberg, 1996a, 2002). Further, precipitation of CaCO3 in the presence of these macromolecules resulted in ACC that was stable to transformation in the dry state for over three months. In contrast, the proteins extracted from the calcitic portion of the sponge spicules induced the formation of calcite crystals. Organisms can also control the transformation of ACC into calcite, as evidenced by the calcification process in sea urchin larvae (Beniash, 1997, 1999; Raz, 2003), or into aragonite as has been observed during the formation of the shells of mollusk larvae (Weiss, 2002).

24.5 Control over crystal orientation Orientated crystal growth is a key feature of many biominerals and must result from orientated nucleation on an organised substrate. An excellent example of orientated crystal growth is provided by aragonitic mollusc shell nacre which comprises stacks of tabular aragonite crystals which are separated by interlamellar and vertical organic sheets (Fig. 24.4). The organic matrix comprises thin layers of -chitin, which is associated with silk fibroin-like proteins and acidic macromolecules (Weiner, 1980, 1984). The c-axes of the crystals are orientated perpendicular to the shell surface while the a and b axes are co-aligned within a given stack. Separate stacks may be aligned or randomly orientated, according to the species of mollusc. The aragonite crystals nucleate at specific sites on the pre-deposited matrix such that the a axis of the aragonite lattice is aligned with the direction of the chitin fibres. The structure of the organic matrix of the mollusc nacreous layer structure is open to discussion (Levi-Kalisman, 2001). Much experimental evidence

24.4 SEM image of a cross-section through an abalone shell showing stacks of orientated aragonite crystals.

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suggests that it comprises thin layers of -chitin, sandwiched between layers of silk fibroin-like proteins in a -sheet conformation and orientated orthogonal to the chitin fibrils, onto which acidic macromolecules were adsorbed. Binding of Ca2+ ions to the orientated proteins was suggested to mimic the arrangement of ions in the ab face of the nascent aragonite crystal, causing the crystal to nucleate with its c axis perpendicular to the nucleating surface. This structural model was principally developed from TEM and X-ray diffraction analyses of dried samples. However, use of cryo-TEM to study the nacreous layer in the hydrated state has suggested that the silk is actually in the form of a hydrated gel, located between, rather than within the sheets of chitin fibrils (LeviKalisman, 2001). The acidic macromolecules may be situated in localised areas on the surfaces of the chitin layers which act as nucleation sites for the aragonite crystals, as well as within the silk gel (Fig. 24.5). This picture of the structure of the organic matrix of mollusc shell nacre therefore contradicts many of the previous ideas on orientated crystal growth in this organism.

24.5 Schematic representation of the structure of the organic matrix of the nacreous layer of the mollusc Atrina serrata, showing interlamellar sheets comprising -chitin and a silk gel located between these sheets. Aspartic acid rich glycoproteins are located as electron dense domains on the surfaces of the interlamellar sheets, and distribute throughout the gel phase. Reproduced with permission from (Levi-Kalisman, 2001).

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A clear demonstration that the combined action of an organised organic matrix and soluble macromolecules can induce both polymorph selectivity and orientated nucleation was obtained on precipitating calcium carbonate on the nucleating protein sheets isolated from the mollusc, in the presence of proteins extracted from the calcitic and aragonitic layers of the mollusc shell (Belcher, 1996). In the absence of soluble proteins, rhombohedral calcite crystals nucleated on the protein sheet, while calcite crystals with spherulitic morphologies were produced in the presence of the calcite-derived proteins. Significantly, the aragonite-derived proteins induced the formation of aragonite crystals, which had needle morphologies and were orientated in the plane of the nucleating protein layer. Aragonite crystals again precipitated when a mixture of the aragonitic and calcitic proteins was used, but were in the form of (001) orientated polycrystalline plates comprising orientated stacks of crystals which were similar in form to molluscan nacre. Growth of calcite seed crystals in solutions of the proteins yielded similar results, with the aragonitic proteins inducing overgrowth of aragonite needles on the calcite crystals and the calcitic proteins causing calcite overgrowth. The nucleating protein sheet therefore dictates the orientation of the calcite primer layer, while the soluble proteins are active in controlling further aspects of shell growth such as crystal polymorph and morphology.

24.6 Control over morphology 24.6.1 Amorphous biominerals Possibly the most striking feature of many biominerals is their remarkable morphologies. Many of these unusual morphologies occur when the biomineral is amorphous, such as amorphous silica in diatoms (Fig. 24.6). An amorphous material has no preferred morphology and thus is readily moulded into the desired shape. The skeleton (frustrule) of diatoms is formed from two interlocking valves, one of which (the epitheca) overlaps the other (the hypotheca) like a lid, together with a number of girdle bands that surround the valves. The top of the frustrule is perforated with many pores, the patterning of which is species-specific, which allow transfer of chemical species between the cell and its environment (Hildebrand, 2003). Diatoms reproduce through cell division. The nucleus first divides and two new valves are formed within the cell wall. The pair of parent valves then separate and fit over the new valves, such that each daughter cell contains a parental epitheca and a new hypotheca. The dividing population of diatoms therefore reduces in size with time, but size is maintained by sexual reproduction (BaÈuerlein, 2003; Li, 1984). Development of the morphology of the skeleton of a diatom can be illustrated by looking at the diatom Coscinodiscus wailesii, in which the silicon skeleton forms in a precisely timed sequence of interactions (Schmid, 1983; Crawford, 1986) (Fig. 24.7). Silicification in all diatoms occurs within a specialised

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24.6 SEM image of a diatom.

24.7 Schematic diagram of valve formation in Coscindiscus wailesii. (a) Initial deposition of silicified strands and formation of areolae in base layer, giving rough structure of silica. (b) Initial smoothing of deposited silica, (c) Final smoothing of base layer (d) Perpendicular growth of areolar walls (e) Growth of silicified `teeth' from the tops of the areolar walls f, f0 , g and g0 . (h) Growth of silica perpendicular to silicified teeth. Reproduced with permission from Schmid (1983).

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vesicle, the silica deposition vesicle (SDV), the shape of which is determined by various components of the cell. An initial pattern for the location of future silica deposition is formed by a close packed array of large vesicles attached to the cell wall (the areolar vesicles) which determine the location of the shell pores and vesicles of the endoplasmic reticulum (ER) which define the location of the basal silica layer. The SDVs then form with tubular morphologies and extend first in the plane of the cell wall between the areolar vesicles, and then vertically between the vesicles. Polymerisation of silicic acid and deposition of silica proceeds continuously with growth of the SDV. Vertical growth stops when the areolar vesicles detach from the cell wall, allowing the SDV to expand laterally and become attached to the cell wall. The vacant pores become filled with small vesicles, which generate a mould for further patterned silica deposition.

24.6.2 Crystalline biominerals While it is easily rationalised how the morphology of amorphous silica can be moulded in this way, it is quite amazing to observe that organisms can similarly produce single crystals with complex shapes and curved surfaces. Crystals typically display sets of planar, low energy faces which are characteristic for a given crystal, and are determined by the crystal structure and symmetry, and the conditions in which the crystal grows. Organisms have developed mechanisms which override this basic growth form to produce crystals whose overall

24.8 SEM image of a sea urchin spine, which is a single crystal of calcite. The inset shows a cross-section through the spine.

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morphologies often bear no relation to the symmetry of the crystal lattice. An excellent example is provided by the skeletal elements of sea urchins (Fig. 24.8). Despite the sponge-like bicontinuous structure and curved surfaces, each skeletal plate and spine is a single crystal of calcite, as demonstrated by X-ray diffraction and polarised light microscopy. In fact, so remarkable are these morphologies, that it was long disputed whether these structures were single crystals or aligned polycrystalline arrays, although analysis by synchrotron Xray diffraction (Berman, 1990) and high-resolution TEM (Su, 2000) has confirmed their single crystal character. A number of mechanisms are believed to be involved in controlling crystal growth. Changes in the activity or positioning of ion pumps and channels during mineralisation may lead to crystal growth in preferred directions. Alternatively, growth may be promoted in specific locations by introduction of storage granules of enriched ions or amorphous material into the mineralisation site (Beniash, 1999). Interaction of growing crystals with soluble additives, in the form of the control macromolecules which are occluded within many biominerals, can also produce subtle changes in morphology. More dramatic changes in morphology are imposed by the physical form of the compartment in which mineralisation occurs. Soluble additives Soluble additives can influence crystal morphology by binding to crystals during growth. Non-specific adsorption is frequently observed, producing particles with rounded surfaces and no well-defined crystal faces. More interesting is when there is a specific interaction between an additive and certain crystal faces only, which is apparent in the formation of new, well-defined crystal faces. Such an effect has been demonstrated by the interaction of a series of ,!dicarboxylates, (structure CO2H±(CH2)n±CO2H)) on the morphology of precipitated calcite crystals (Mann, 1990). The effect of the molecules depended on the separation of the carboxylate groups, with malonic acid (n ˆ 1) being the most effective in inducing a morphological change. At a Ca/additive ratio of  3, crystals were elongated parallel to the c-axis, were capped with rhombohedral {104} end faces, and displayed curved {110} faces approximately parallel to the axis of elongation. Similar morphologies were observed with other dicarboxylate molecules, but this effect was reduced with increase in the chain length. These morphological changes were considered to arise from molecular recognition between specific crystal faces and the additive, resulting in the appearance of a set of symmetry-related faces. The crystal faces of calcite lying parallel to the c-axis contain carbonate groups orientated perpendicular to the face. The dicarboxylate additives present a bidentate carboxylate group which can mimic the presentation of carboxylate groups in the crystal. These crystal

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faces could therefore be stabilised via stereoselective adsorption of the acids. The reduced effect of the dicarboxylic acids with increasing chain lengths can be attributed to cooperative binding of the carboxylate groups in the short-chain additives but the carboxylate groups acting independently in the longer-chain additives. Experiment has demonstrated that the soluble macromolecules extracted from within calcium carbonate biominerals can also adsorb to specific calcite crystal faces during precipitation, resulting in modified crystal morphologies (Albeck, 1993, 1996b; Aizenberg, 1994, 1995a). The macromolecules are postulated to possess structural motifs that match the atomic arrangement on one set of crystal planes, causing the additive to interact with and stabilise these faces (Addadi, 1985). Following adsorption to specific crystal faces, the macromolecules become overgrown and are occluded within the crystals, as demonstrated by calcite crystals grown in the presence of fluorescence-labelled sea urchin macromolecules (Berman, 1988). The occluded macromolecules are unique to a particular biomineral. This can be exemplified by the action of macromolecules extracted from spines of the sea urchin Paracentrotus lividus (Albeck, 1996b). These proteins were chemically and enzymatically treated to yield three fractions; the polysaccharides were removed yielding deglycosylated protein, whole polysaccharide chains were isolated, and densely glycosylated peptide cores were produced. The whole assembly of macromolecules interacted with crystal planes approximately parallel to the c-axis of calcite, producing faces indexed as (10l) where l  1±5, but which are not well-defined. The non-specific activity of the isolated polysaccharide fraction can be attributed to loss of ordered conformation on release into solution. Conformational stability is typically achieved when a large number of polysaccharides act cooperatively, such as occurs when several chains are located on a protein core. The glycosylated proteins interacted with well-defined crystal faces, suggesting that densely glycosylated regions of the peptides form assemblies of ordered structure. Therefore, while the densely glycosylated proteins and the deglycosylated proteins interact specifically with one set of calcite planes, the whole assembly of macromolecules do not, inducing instead the formation of several closely related faces oblique and parallel to the c-axis. The polysaccharide groups on the protein therefore contribute to a range of interactions that are not observed on removal of the majority of polysaccharide from the glycoproteins. Direct evidence of the interaction of biological macromolecules with the faces of growing crystals has been obtained using AFM (Walters, 1997). Imaging of the {104} faces of a rhombohedral calcite crystal in saturated calcium carbonate solution shows straight step edges with some kinks, and sharp corners between different step edges. Brief exposure to intra-crystalline proteins extracted from the calcitic layers of the red abalone Haliotis rufescens resulted in rounding of the corners between step edges, and the step edges appeared more

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convoluted and kinked. The proteins agglomerated at the step edges and appeared to attach more strongly to step edges than to the terraces. Proteins extracted from the aragonitic layer of the abalone behaved quite differently, causing step edges to become highly convoluted, terraces extremely rough, and finally terminated crystal growth. A change in the lattice structure of the calcite {104} face was also apparent, and was consistent with overgrowth of aragonite from a (001) plane. It has also been suggested that organic additives influence crystal morphologies by binding to the surface-step edges rather than single crystal faces. In this way, the crystal morphologies produced depend on both stereochemical recognition and the effects of binding on the interfacial energies of the growing crystal (Orme, 2001). AFM study of the growth of the calcite (10.4) face in the presence of the chiral molecules D-aspartic acid and L-aspartic acid showed that interaction with the crystal steps was asymmetric. Modelling of the interaction of these molecules with the step edges suggested that D-Asp binds to the (014) riser such that one of its negatively charged carboxyl groups completes the coordination of calcium ions, while the positively charged (NH 3+) group remains in registry with the positive ions at the surface. At the step, the remaining carboxyl group can bind to the two adjacent {10.4} surfaces. In contrast, the carboxyl group of D-Asp does not closely match the carbonate groups on the (114) riser. The specific amino acid enantiomers therefore bind to the step edges that offer the best geometric and chemical fit. This changes the step-edge free energies, which in turn results in macroscopic crystal shape modifications. Spatial constraint While control macromolecules are clearly active in modifying morphologies, their action alone cannot be responsible for producing crystals with shapes as complex as those of the sea urchin skeletal elements. In common with amorphous materials, crystals with unusual morphologies and curved surfaces are typical of growth within vesicles. Many biominerals, such as the calcite scales produced by coccoliths (Young, 2003), or sea urchin larval spicules (Beniash, 1997), form within vesicles with defined shapes and grow until they impinge upon the vesicle, which effectively acts as a mould. The size and shape of the vesicle may be altered during the crystal growth process. However, while it would be anticipated that an amorphous material will take up the form of a soft organic membrane, it is perhaps surprising that it can impose a form on a crystal whose energetically favoured form comprises well-defined faces. Interaction of the crystal with the vesicle membrane may stabilise the high-energy rounded crystal surfaces. Sea urchin larvae may actually employ an amorphous calcium carbonate (ACC) precursor phase to assist control of morphology (Beniash, 1997, 1999; Raz, 2003). Larval spicules initially present a tri-radiate morphology and curved

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surfaces, and yet are single crystals of calcite. Mineralisation is intravesicular and begins by formation of an orientated, regular rhombohedral crystal of calcite. Growth then continues by addition of ACC, and the mineral deposit gradually transforms into a single crystal of calcite with non-crystallographic morphology. ACC is present in relatively large amounts during the first stages of spicule growth and decreases with ageing of the spicule. The transformation process of the biogenic ACC is much slower than for synthetic ACC but is also significantly faster than the crystallisation of dried spicules isolated from the larvae, suggesting that the organism actively controls crystallisation. Experiments investigating crystallisation of ACC within a constrained volume have supported the suggestion that an amorphous precursor phase can aid morphological control of crystals. Calcium carbonate was precipitated in the 3 m diameter, cylindrical pores of track-etch membranes at room and low temperature (4±6 ëC) (Loste, 2001). At low temperatures, the intra-membrane particles were rods of dimensions 3 m  10 m, which had clearly been dictated by the geometry of the membrane channels. 0.5 m spherical particles of ACC initially coated the walls of the pores before filling in the entire volume and generating the final cylindrical form (Fig. 24.9(a)). Transformation to single crystals of calcite occurred with time, while maintaining the rod-like morphology (Fig. 24.9(b)). At room temperature, when calcite is precipitated directly rather than via an ACC phase, crystals with irregular morphologies were produced. These results therefore suggest that transformation of ACC within a constrained volume can therefore produce calcite crystals of morphology imposed by the environment. Single crystals of calcite with complex form have also been produced in the absence of additives by precipitation in a polymer membrane with sponge-like structure (Park, 2002, 2004). The membrane was templated by a sea urchin plate and was produced by dipping a plate in the polymer monomer solution, curing the polymer, and finally dissolving away the calcium carbonate to generate the polymer replica. As the porous and inorganic fractions of the sea urchin plate occupy equal volumes, and have identical morphologies, the polymer membrane produced had an identical structure to the original calcium carbonate plate. Calcium carbonate was precipitated in the membrane by the double diffusion technique, with the product depending on the solution concentrations. At higher solution concentrations, a limited network of polycrystalline calcite crystals was formed (Fig. 24.10(a)). In contrast, at low reagent concentrations, large single crystals of calcite with morphologies identical to the structure of the polymer membrane were produced (Fig. 24.10(b)). These particles showed both non-crystallographic curved surfaces, as well as the planar faces characteristic of crystalline materials. The results demonstrate that single crystals with intricate structures can be produced synthetically, in the absence of additives, by precipitation in a suitable constrained volume.

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24.9 Calcium carbonate particles precipitated in the 3 m pores of track etch membranes isolated after (a) 15 minutes, showing the construction from spherical amorphous calcium carbonate particles and (b) 24 hours, showing uniform structure.

Large sheets of patterned single crystal calcite have also been formed by crystallisation of ACC in a restricted volume (Aizenberg, 2003). In this case, the template comprised a square array of posts, sandwiched between two substrates. One substrate was coated with a thin gold film, and was functionalised with a mixture of phosphate-, methyl- and hydroxyl-terminated thiols. A single nucleation centre comprising an area of hydroxyl, carboxyl or sulphonate terminated thiols was introduced onto the substrate using an AFM tip. The second substrate was a gas-permeable polymer film. Immersion of the prepared template in a solution of calcium chloride and exposure to an atmosphere of CO2 released from solid ammonium carbonate, resulted in initial precipitation of ACC. Nucleation of a calcite crystal at the designated nucleation centre occurred

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24.10 Calcium carbonate precipitated in sponge-like polymer membranes (a) from 0.4 M CaCl2/Na2CO3 showing polycrystalline structure and (b) from 0.02 M CaCl2/Na2CO3 showing single crystal structure.

within approximately one hour, and growth proceeded within the boundaries presented by the network of posts, producing patterned single crystals of calcite up to 1 mm in size. The crystals were specifically orientated with respect to the substrate according to the functional group expressed by the SAM nucleation centre. The induction time for nucleation increased with a decrease in the feature sizes, varying from one hour for an unpatterned substrate, to four to five hours for 1 m channels. Additionally, micropatterned single crystals were not formed in substrates with large channels, but formed crystals 15 to 25 m in size when the nucleated at distances > 15 m from the channel edge. It was suggested that the micropattern may provide a `microsump action' for the release of excess water during the ACC to calcite transition, and/or to provide relaxation of stress in the forming crystal. The crystallisation process may thus occur by mass

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transport between the amorphous and crystalline phases, rather than by a solid state transformation. It is notable that all large biological single crystals are porous. This may be for the reasons postulated above, as growth of large single crystals from an amorphous precursor phase may be associated with high mechanical stresses caused by release of water. In vivo, the organic structures delineating the mineralisation compartment may provide sites for stress relaxation. Formation of multi-crystalline arrays Many biominerals with complex morphologies are polycrystalline structures generated by organised assembly of single crystal components. An excellent example of this is the sphere of calcite scales (coccoliths) formed by the single cell algae Emiliania huxleyi (Fig. 24.11). The coccolith scales each comprise about 30±40 units which are organised in a ring to give a double-rimmed structure. Each unit is a single crystal of calcite, comprising a base plate (the proximal shield element) and a hammer-head shaped upper plate (the distal shield element) which are separated by a central wall. The calcite c-axis lies parallel to the direction of elongation of the elements and in the plane of the

24.11 SEM image of the of the alga Emiliania huxleyi showing the assembly of coccolith plates forming the coccosphere.

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proximal shield element, while the a-axis lies perpendicular to the plane of the ring. The units are then organised in the coccolith such that the c-axis is orientated at 20ë to ring radius. This orientation is always in a clockwise direction, endowing the coccolith with morphological chirality. The orientation of these units must be defined at nucleation (Young, 1999). Formation of coccoliths is intravesicular and begins with formation of an organic base plate scale and assembly of vesicles along the rim of this plate. Nucleation then occurs within the vesicles to generate a proto-coccolith ring of interlinked calcite crystals. The crystals initially form as 40 nm thick rhombohedral plates which are inclined to the plane of the ring. The plates grow to a height of 100 nm, and radial outgrowth along the c-axis from the top and bottom faces generate a Z-shape. These units become inter-linked by selective growth along the inside rim, and further radial growths from the base and top of the element produce the proximal and distal shield elements. The coccolith vesicle is in close contact with the developing crystal during growth, and is likely to play a role in morphological control of the calcite units. Finally, when the coccolith plate has fully developed it moves to the cell wall where it is extruded from the cell to form a spherical shell of coccoliths.

24.7 Control over mechanical properties Many biominerals also possess impressive mechanical properties, which principally derive from their composite character and structural organisation. Mollusc nacre is a polycrystalline calcium carbonate biomineral with excellent mechanical properties. Despite having a very low organic content (1%), this material is superior to most other composite ceramics in stiffness, strength and toughness (Jackson, 1988, 1990) and 3000 times more resistant to fracture than a single crystal of pure aragonite (Currey, 1977). Nacre comprises layers of interlocking aragonite platelets separated by a thin layer of organic material. When a crack propagates through nacre it passes around the platelets by a tortuous path (Wang, 1995; Jackson, 1990) causing the plates to spring apart, and extending the organic sheets (Wang, 1995; Jackson, 1988). The organic `adhesive' between the plates appears to be key in the fracture resistance of this material (Currey, 2001; Smith, 1999), and elongates by pulling open folded domains and loops, or breaking inter-chain bonds (Smith, 1999). Biogenic single crystals can also exhibit mechanical properties far superior to their synthetic counterparts. Considering the skeletal elements of echinoderms, these single crystals of calcite have strength-to-weight ratios exceeding those of many man-made construction materials such as brick and concrete (Weber, 1969; Currey, 1975) and fracture to give conchoidal fracture surfaces. Despite being single crystals, sea urchin calcite is actually a composite material, occluding 0.1 wt% organic macromolecules which interfere with crack propagation along the cleavage planes. Incorporation of macromolecules within

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biological single crystals has been studied using high-resolution synchrotron Xray diffraction (XRD) (Aizenberg, 1995a, 1995b, 1996b, 1997). Occlusion of macromolecules within single crystals causes defects which can be described in terms of the coherence length (the size of perfect crystalline domains) and the domain spread (the misalignment of the crystalline blocks). Measurement of these factors along different crystallographic directions provides information on the location of the macromolecules. A series of single crystal, calcite sponge spicules with different morphologies was analysed to investigate the textures of the crystals. Measurement of coherence lengths along different crystallographic directions enabled the average shape of domains to be determined, and revealed a distinct correlation with the overall morphology of the spicules. Generally, longer coherence lengths were present in crystallographic directions close to the morphological axes of the spicules. This was true only for spicules containing occluded macromolecules and no correlation was detected for pure calcite spicules (Aizenberg, 1995a). In the case of spicules with morphologies which did not reflect the symmetry of the calcite lattice, coherence lengths of symmetry-related faces were not identical (Aizenberg, 1995a). This suggests that control beyond specific protein-crystal interactions must be operative, creating an anisotropic environment for crystal growth (Aizenberg, 1996b). This could be achieved through orientated nucleation of a crystal from a specific crystallographic plane, such that it grew in one direction only. Macromolecules could then be adsorbed on specific planes that are exposed during growth.

24.8 Conclusion Biological systems are capable of controlling mineralisation processes to a remarkable degree, and in the absence of the high temperatures and pressures available in synthetic experiments, principally achieve this using organic molecules as soluble additives and insoluble frameworks. Although isolation and analysis of the key macromolecules associated with controlling mineral formation is extremely challenging, recent work (Gotliv, 2003) has offered a novel route to isolation and fractionation of macromolecules from calcium carbonate biominerals. It is envisaged that further research in this area will enable sequencing of the macromolecules used to select, for example, calcite or aragonite, and thus permit development of synthetic analogues. Significant recent advances in the understanding of biomineralisation control strategies include the demonstration that in certain organisms, calcite and aragonite formation can occur via an amorphous calcium carbonate (ACC) precursor. Combined with the observation that ACC appears to contain a degree of short-range structure which defines the structure of the product phase subsequent to crystallisation, it is clear that amorphous phases may provide a route to novel control of material structure. The observation that mollusc nacre

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may develop in a gel-like environment also provides new insight into the formation of hierarchically organised mineral structures. These, and other mechanisms used in biological mineralisation provide inspiration for control of the structure and properties of minerals. Future research will ultimately lead to the preparation of functional materials based on these ideas.

24.9 Further information A number of very good review articles and books, which have been cited in this chapter, provide a more detailed insight into the subject of biomineralisation. For a general overview of the topic of biomineralisation, On Biomineralization by Lowenstam HA and Weiner S is highly recommended, providing a very accessible and thorough examination of biomineralisation in all five kingdoms, while Mann S (2001), in Biomineralization: Principles and Concepts in Bioinorganic Materials Chemistry gives an introduction to the principles of biomineralisation and their application to biomimetic materials chemistry. More specialised recent review articles discuss biomineralisation in selected organisms considered in this chapter; magnetotactic bacteria in Bazylinski (2004), coccolith formation in Young (2003) and silica biomineralisation in diatoms and sponges in MuÈller (2003). The role of amorphous calcium carbonate in biology is reviewed in Addadi (2003) while Meldrum (2003) provides a review of the control strategies used by organisms during calcium carbonate biomineralisation, and their subsequent application to biomimetic engineering of calcium carbonate in synthetic systems.

24.10 References Addadi L and Weiner S (1985) `Interactions between acidic proteins and crystals: Stereochemical requirements in biomineralization' Proc. Natl. Acad. Sci., 82(12), 4110±4114. Addadi L, Moradian J, Shay E, Maroudas NG and Weiner S (1987) `A chemical model for the cooperation of sulfates and carboxylates in calcite crystal nucleation: Relevance to biomineralization' Proc. Natl. Acad. Sci. USA, 84(9), 2732±2736. Addadi L, Raz S and Weiner S (2003) `Taking advantage of disorder: amorphous calcium carbonate and its role in biomineralization' Adv Mater, 15(12), 959±970. Aizenberg J, Albeck S, Weiner S and Addadi L (1994) `Crystal protein interactions studied by overgrowth of calcite on biogenic skeletal elements' J. Crys Growth, 142(1±2), 156±164. Aizenberg J, Hanson J, Ilan M, Leiserowitz L, Weiner S and Addadi L (1995a) `Morphogenesis of calcitic sponge spicules ± a role for specialized proteins interacting with growing crystals' FASEB J. 9(2), 262±268. Aizenberg J, Hanson J, Koetzle TF, Leiserowitz L, Weiner S and Addadi L (1995b) `Biologically induced reduction in symmetry ± a study of crystal texture of clacitic sponge spicules' Chem. Eur. J. 1(7): 414±422. Aizenberg J, Lambert G, Addadi L and Weiner S (1996a) `Stabilization of amorphous

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calcium carbonate by specialized macromolecules in biological and synthetic precipitates' Adv Mater 8(3), 222±226. Aizenberg J, Ilan M, Weiner S and Addadi L (1996b) `Intracrystalline macromolecules are involved in the morphogenesis of calcitic sponge spicules' Connec. Tiss. Res. 35(1±4), 17±23. Aizenberg J, Hanson J, Koetzle TF, Weiner S and Addadi L (1997) `Control of macromolecule distribution within synthetic and biogenic single calcite crystals' J. Am Chem Soc 119(5), 881±886. Aizenberg J, Lambert G, Weiner S and Addadi L (2002) `Factors involved in the formation of amorphous and crystalline calcium carbonate: A study of an ascidian skeleton' J Am Chem Soc, 124(1), 32±39. Aizenberg J, Muller DA, Grazul JL and Hamann DR (2003) `Direct fabrication of large micropatterned single crystals' Science, 299 (5610), 1205±1208. Albeck S, Aizenberg J, Addadi L and Weiner S (1993) `Interactions of various skeletal intracrystalline components with calcite crystals' J Am Chem Soc, 115(15), 11691± 11697. Albeck S, Addadi L and Weiner S (1996a) `Regulation of calcite crystal morphology by intracrystalline acidic proteins and glycoproteins' Connec. Tiss. Res. 35(1-4), 365± 370. Albeck S, Aizenberg J, Addadi L and Weiner S (1996b) `Polysaccharides of intracrystalline glycoproteins modulate calcite crystal growth in vitro' Chem Eur J, 2(3), 278±284. BaÈuerlein E (2003) `Biomineralization of unicellular organisms: an unusual membrane biochemistry for the production of inorganic nano- and microstructures', Angew Chem Int Ed, 42(6), 614±641. Bazylinski DA and Frankel RB (2003) `Biologically controlled mineralization in prokaryotes' in Dove PM, De Toreo JJ and Weiner S Reviews in mineralogy and geochemistry Vol 54, Washington, Mineral Soc Am, pp 217±247. Bazylinski DA and Frankel RB (2004) `Magnetosome formation in prokaryotes', Nature Rev Microbiol 2(3), 217±230. Belcher AM, Wu XH, Christensen RJ, Hansma PK, Stucky GD and Morse DE (1996) `Control of crystal phase switching and orientation by soluble mollusc-shell proteins' Nature 381(6577), 56±58. Beniash E, Aizenberg J, Addadi L and Weiner S (1997) `Amorphous calcium carbonate transforms into calcite during sea urchin larval spicule growth' Proc. Roy. Soc. Lon. B, 264 (1380): 461±465. Beniash E, Addadi L and Weiner S (1999) `Cellular control over spicule formation in sea urchin embryos: A structural approach' J Struct Biol, 125 (1): 50±62. Berman A, Addadi L and Weiner S (1988) `Interactions of sea-urchin skeleton macromolecules with growing calcite crystals ± a study of intracrystalline proteins' Nature, 331(6156), 546±548. Berman A, Addadi L, Kvick A, Leiserowitz L, Nelson M and Weiner S (1990) `Intercalation of sea-urchin proteins in calcite ± study of a crystalline composite material' Science, 250 (4981), 664±667. BruÈmmer F (2003) `Living inside a glass box ± silica in diatoms' in MuÈller WEG, Silicon biomineralization: biology, biochemistry, molecular biology, biotechnology, Progress in Molecular and subcellular biology Vol 33, Berlin, Springer, pp 3±10. Cha JN, Shimizu K, Zhou Y, Christiansen SN, Chemlka BF, Stucky GD and Morse DE

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(1999) `Silicatein filaments and subunits from a marine sponge direct the polymerization of silica and silicones in vitro' Proc. Natl. Acad. Sci., 96(2), 361± 365. Chasteen ND and Harrison PM (1999) `Mineralization in ferritin: an efficient means of iron storage', J. Struct. Biol., 126(3), 182±194. Crawford RM and Schmid AM (1986) `Ultrastructure of silica deposition in diatoms' in Leadbeater BS and Riding R Biomineralization in lower plants and animals, The Systematics Association Special Volume Number 30, Oxford, Oxford University Press, pp 291±314. Currey JD (1975) `Comparison of strength of echinoderm spines and mollusk shells' J Mar Biol Assoc U.K 55, 419±424. Currey JD, (1977) `Mechanical properties of mother-of-pearl in tension', Proc. R. Soc. Lond. B, 1977, 196(1125), 443±463. Currey JD, Zioupos P, Davies P and Casinos AJ (2001) `Mechanical properties of nacre and highly mineralized bone' Proc. R. Soc. Lond. B, 268(1462), 107±111. Falini G, Albeck S, Weiner S and Addadi L (1996) `Control of aragonite or calcite polymorphism by mollusk shell macromolecules' Science 271(5245), 67±69. Frankel RB, Papaefthymiou GC, Blakemore RP and O'Brien W (1983) `Fe3O4 precipitation in magnetotactic bacteria' Biochim. Biophys. Acta 763(2), 147±159. Gotliv BA, Addadi L and Weiner S (2003) `Mollusk shell acidic proteins: In search of individual functions' ChemBioChem, 4(6), 522±529. Harrison PM and Arosio P (1996) `Ferritins: Molecular properties, iron storage function and cellular regulation', Biochim. Biophys. Acta Bioenerg 1275(3), 161±203. Hildebrand M and Wetherbee R (2003) `Components and control of silicification in diatoms' in MuÈller WEG, Silicon biomineralization: biology, biochemistry, molecular biology, biotechnology , Progress in Molecular and subcellular biology Vol 33, Berlin, Springer, pp 11±57. Hunter GK (1996), `Interfacial aspects of biomineralization' Curr. Op. Solid State and Int. Sci., 1(3), 430±435. Jackson AP, Vincent JFV and Turner RM (1988) `The mechanical design of nacre' Proc. R. Soc. Lond. B, 234(1277), 415±440. Jackson AP, Vincent JFV and Turner RM (1990) `Comparison of nacre with other ceramic composites' J. Mater. Sci., 1990, 25(7), 3173±3178. Komeili A, Vali H, Beveridge TJ and Newman DK (2004) `Magnetosome vesicles are present before magnetite formation, and MamA is required for their activation' Proc Nat Acad Sci, 101(11), 3839±3844. KroÈger N, Deutzmann and Sumper M (1999) `Polycation peptides from diatom biosilica that direct silica nanosphere formation' Science, 286(5442), 1129±1132. KroÈger N, Deutzmann R, Bergsdorf C and Sumper M (2000) `Species-specific polyamines from diatoms control silica morphology', Proc. Natl. Acad. Sci., 97(26) 14133±14138. KroÈger N, Deutzmann R, Bergsdorf C and Sumper M (2001) `Silica-precipitating peptides from diatoms', J. Biol. Chem. 276(28), 26066±26070. KroÈger N, Lorenz S, Brunner E and Sumper M (2002) `Self-assembly of highly phosphorylated silaffins and their function in biosilica morphogenesis' Science, 298 (5593), 584±586. Levi Y, Albeck S, Brack A, Weiner S and Addadi L (1998) `Control over aragonite crystal nucleation and growth: An in vitro study of biomineralization' Chem Eur J,

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4(3), 389±396. Levi-Kalisman Y, Falini G, Addadi L and Weiner S (2001) `Structure of the nacreous organic matrix of a bivalve mollusk shell examined in the hydrated state using CryoTEM' J. Struct. Biol. 135(1), 8±17. Li CW and Volcani BE (1984) `Aspects of silicification in wall morphogenisis of diatoms' Phil Trans Roy Soc Lon B, 304(1121), 519±528. Lippmann F (1973) Sedimentary Carbonate Minerals, Berlin, Springer-Verlag. Loste E and Meldrum FC (2001) `Control of calcium carbonate morphology by transformation of an amorphous precursor in a constrained volume' Chem. Commun. 10, 901±902. Lowenstam HA and Weiner S (1989), On Biomineralization, New York, OUP. Mann S, Frankel RB and Blakemore RP (1984) `Structure, morphology and crystal growth of bacterial magnetite' Nature 310(5976), 405±407. Mann S, Didymus JM, Sanderson NP, Heywood BR, Samper EJA (1990) `Morpholgoical influence of functionalized and non-functionalized-alpha, omega-dicarboxylates on calcite crystallization' J. Chem. Soc. Faraday Trans, 86(10), 1873±1880. Mann S (2001), Biomineralization: Principles and Concepts in Bioinorganic Materials Chemistry, Oxford, OUP. Meldrum FC, Mann S, Heywood BR, Frankel RB and Bazylinski DA (1993) `Electron microscopy study of magnetosomes in two cultured vibrioid magnetotactic bacteria' Proc R Soc Lon B, 251(1332), 237±242. Meldrum FC (2003) `Calcium carbonate in biomineralisation and biomimetic chemistry' Int Mater Rev, 48(3), 187±224. MuÈller WEG ed., (2003) Silicon biomineralization: biology, biochemistry, molecular biology, biotechnology, Progress in molecular and subcellular biology Vol 33, Berlin, Springer. Orme CA, Noy A, Wierzbicki A, McBride MT, Grantham M, Teng HH, Dove PM, DeYoreo JJ (2001) `Formation of chiral morphologies through selective binding of amino acids to calcite surface steps' Nature, 411(6839), 775±779. Park RJ and Meldrum FC (2002) `Synthesis of single crystals of calcite with complex morphologies' Adv Mater,14 (16), 1167±1169. Park RJ and Meldrum FC (2004) `Shape-constraint as a route to calcite single crystals with complex morphologies' J Mater Chem, in press. Perry CC and Keeling-Tucker T (2003) `Model studies of colloidal silica precipitation using biosilica extracts from Equisetum telmateia' Coll. Polym. Sci. 281(7), 652± 664. Poulsen N, Sumper M, KroÈger N (2003) `Biosilica formation in diatoms: Characterization of native silaffin-2 and its role in silica morphogenesis' Proc Natl Acad Sci USA, 100(21), 12075±12080. Raz S, Hamilton PC, Wilt FH, Weiner S and Addadi L (2003) `The transient phase of amorphous calcium carbonate in sea urchin larval spicules: The involvement of proteins and magnesium ions in its formation and stabilization' Adv Func Mater, 13 (6): 480±486. Schmid A-MM and Volcani BE (1983) `Wall morphogenesis in Coscinodiscus wailesii. I. Valve morphology and development of its architecture' J. Phycol 19(4), 387±402. SchuÈler D and BaÈuerlein E (1998) `Dynamics of iron uptake and Fe3O4 mineralization during aerobic and microaerobic growth of Magnetospirillum gryphiswaldense AMB-1' J. Bacteriol. 180(1), 159±162.

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SchuÈ ler D and Frankel RB (1999) `Bacterial magnetosomes: microbiology, biomineralization and biotechnological applications' Appl. Microbiol. Biotechnol. 52(4), 464±473. Shimizu K, Cha JN, Stucky GD and Morse DE (1998), `Silicatein a: cathepsin L-like protein in sponge biosilica' Proc. Natl. Acad. Sci., 95(11), 6234±6238. Smith BL, Schaffer TE, Viani M, Thompson JB, Frederick NA, Kindt J, Belcher A, Stucky GD, Morse DE, and Hansma PK (1999) `Molecular mechanistic origin of the toughness of natural adhesives, fibres and composites' Nature, 399 (6738): 761±763. Su X, Kamat S and Heuer AH (2000), `The structure of sea urchin spines, large biogenic single crystals of calcite' J. Mat Sci, 35 (22), 5545±5551. Sumper M, Lorenz S and Brunner E (2003) `Biomimetic control of size in the polyaminedirected formation of silica nanospheres' Angew Chem Int Ed 42(42), 5192±5195. Walters DA, Smith BL, Belcher AM, Paloczi GT, Stucky GD, Morse DE and Hansma PK (1997) `Modification of calcite crystal growth by abalone shell proteins: An atomic force microscope study' Biophys J., 72(3), 1425±1433. Wang RZ, Wen HB, Cui FZ, Zhang HB and Li HD (1995) `Observations of damage morphologies in nacre during deformation and fracture' J. Mater. Sci., 30(9), 2299± 2304. Weber J, Greer R, Voight B, White E and Roy R, (1969) `Unusual strength properties of echinoderm calcite related to structure' J. Ultrstruc. Res. 26, 355±366. Weiner S and Hood L (1975) `Soluble protein of the organic matrix of mollusk shells: a potential template for shell formation' Science, 190(4218), 987±989. Weiner S and Traub W (1980) `X-ray diffraction study of the insoluble organic matrix of mollusk shells' FEBS Letts, 111(2), 311±316. Weiner S and Traub W (1984) `Macromolecules in mollusk shells amd their functions in biomineralization' Phil Trans Roy Soc Lon B, 304(1121), 425±434. Weiner S and Addadi L (1997) `Design strategies in mineralized biological materials' J Mater Chem, 7(5), 689±702. Weiss IM, Tuross N, Addadi L and Weiner S (2002) `Mollusc larval shell formation: Amorphous calcium carbonate is a precursor phase for aragonite' J Exp Zool 293(5), 478±491. Wheeler AP, George JW and Evans CA (1981) `Control of calcium carbonate nucleation and crystal growth by soluble matrix of oyster shell' Science, 212(4501), 1397± 1398. Wheeler AP and Sikes CS (1984) `Regulation of carbonate calcification by organic matrix' Amer Zool, 24(4), 933±944. Wheeler AP, Low KC and Sikes CS (1991) `CaCO3 crystal-binding properties of peptides and their influence on crystal growth' in Sikes CS and Wheeler AP Surface Reactive Peptides and Polymers Vol. 444, Washington, ACS, 72±84. Young JR, Davis SA, Brown PR and Mann S (1999) `Coccolith ultrastructure and biomineralization' J. Struct. Biol. 126(3), 195±215. Young JR, Henriksen K (2003) `Biomineralization within vesicles: The calcite of coccoliths' in Weiner S and Dove PM Biomineralization. Reviews in mineralogy and geochemistry 54, Washington, Mineral. Soc. America, pp 189±215.

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P E T O M L I N S and R L E A C H , National Physical Laboratory, UK, P V A D G A M A , University of London, UK, S M I K H A L O V S K Y and S J A M E S , University of Brighton, UK

25.1 Introduction It has long been recognised that the surface chemistry and surface texture of biomaterials that are in contact with body fluids are critically important for maximising biocompatibility. The tendency of cells to adhere or not to adhere to materials is essential for their performance, for example, in tissue scaffolds or urinary stents. A wealth of literature deals with the development of coatings or finishes that are designed to `control' or influence cellular responses, yet few publications attempt to quantify the surface texture of these materials. This chapter provides an overview of surface texture assessment in terms both of the parameters used and the surface texture measurement methods that are available. Additional, more detailed information for specific techniques can be found in the texts cited in section 25.12. The literature concerned with surface texture measurement is confusing; it contains many and varied descriptions of surface texture parameters. Part of this confusion can be attributed to a change in the notation used to represent texture between the 1984 and 1997 versions of the international standard, ISO 4287 that defines surface roughness terminology. The terms and definitions used in this chapter are consistent with ISO 4287: 1997.

25.2 Biomaterials, surfaces and biocompatibility A biomaterial can be defined as a solid substance (other than drugs) or a combination of synthetic or natural substances that can be used for any period of time, as a whole or as a part of a system that treats, augments, or replaces any tissue, organ, or function of the body. This broad definition covers metals, ceramics and polymers that are used in applications such as tissue-engineering scaffolds, catheters, pins, screws, drug delivery housings, stents and substitute materials, e.g., bone and skin, and include fluids although these materials are not relevant for this chapter. Medical implants and devices that come into contact with body fluids are manufactured from biomaterials using processing techniques that include casting,

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moulding, extrusion and machining. The surfaces of these objects are not completely smooth, in the literal sense, or devoid of irregularities and may have characteristic micro-scale features akin to a fingerprint that reflect the surface features of the mould cavity or machining process. The surface of a biomaterial plays a key role in determining biocompatibility influencing, for example, both cell adhesion (Bowers et al., 1992, Keller et al., 1994, Rosa et al., 2003, Hallab et al., 2001) and gene expression (Schneider et al., 2003, Ogawa et al., 2002).

25.3 What is a surface? A surface is defined according to ISO 4287: 1997 as something that limits a body and separates it from the surrounding medium. Surface texture can be defined as any repetitive or random departure from a nominal surface. A twodimensional profile of a surface is produced at the intersection of a plane drawn perpendicular to it. The orientation of the profiles selected to represent a threedimensional surface can have a marked effect on the parameters that are used to describe the texture, especially for surfaces that have a well-defined pattern. This problem can be surmounted by defining the surface profile according to ISO 13565-1: 1996 which states that the traversing direction for assessment purposes should be perpendicular to the direction of lay unless otherwise stated. Lay is defined as the direction of a dominant surface pattern. In terms of coordinates, y is defined as being parallel with the lay direction and x perpendicular to it, z is out-of-plane. Some surfaces will have a well-defined texture with no preferred orientation, for these cases several two-dimensional profiles at different orientations should be measured and the maximum value taken as a measure of the roughness. The profile of surfaces that have a uniform, nondirectional texture can be taken at any in-plane orientation.

25.4 Surface measurement Surface texture is of paramount importance in engineering especially for applications where surfaces come into contact with each other. Techniques such as grinding, shot peening and lapping produce characteristic, often uniform, surfaces in metals that lend themselves to parametric characterisation. Hence this area of metrology has a strong engineering influence. Most surface texture measurements are made using stylus-based devices in direct contact with the surface where the movement of a probe is monitored as it traverses over it. These instruments are not suitable for measuring the texture of soft materials, as they tend to damage the surface. Non-contacting instruments that rely on optical measurements or current flow are more appropriate methods for obtaining data from these materials. Conventional stylus instruments measure the surface profile over one or more sampling lengths (Leach, 2001). Typically five sequential sampling lengths are

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used (ISO 4287: 1997) which combine to form the evaluation length, ln for quantification of roughness. (Standard procedures have yet to be developed for laser profilometers and atomic force microscopes.)

25.5 Filters Surface data can be filtered to remove unwanted noise or to remove texture information at unwanted wavelengths. Filters are classified according to the spatial periodicity that they allow to pass through; low-pass filters admit long wavelengths and reject short ones; high-pass filters do the opposite. Band-pass filters, as the name implies, allow a limited range of wavelengths to pass. In practice using filters can create problems in deciding how much of the noise in the measurements is `real' and how much can be attributed to the surface. Filters used in surface texture measurements do not have a sharp cut-off in frequency above or below which information is rejected. This gradual attenuation of high or low frequency data helps avoid distortion of the measurements as can occur when strong features are close to the filtration limits. The point on the transmission curve at which the transmitted signal is reduced to 50% is referred to as the cut-off wavelength, c, of the filter (Fig. 25.1). The choice of c depends on the sampling frequency and the speed of the instrument, for example, measurements made at intervals of 0.01 mm from a device moving at 1 mmsÿ1 will generate data at a frequency of 100 Hz. Increasing the sampling interval to 0.1 mm will reduce the frequency at which data are obtained to 10 Hz. A high pass filter that suppresses all frequencies below 10 Hz effectively

25.1 A schematic representation of a low-pass filter.

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removes any surface irregularities larger than 0.1 mm spacing from the data. Hence filters can be used to bias the experimental data towards detecting profile (surface texture after applying a low pass filter to the data), waviness (after applying a band pass filter) and roughness (after applying a high pass filter). Measurement conditions are set according to ISO 3274: 1996 for Gaussian filters according to the respective values of the sampling interval, measurement speed and filtration limits.

25.6 Quantifying surface texture Before measured surface texture profiles in two dimensions or threedimensional areal data can be assessed or compared, the experimental data have to be quantified, a process that usually involves reducing large amounts of information into a small set of parameters. This process is by no means straightforward, many of the parameters in common usage, e.g., the arithmetic mean, `Ra' give the impression that the surface is smoother than it actually is. International standards are available for quantifying measured two-dimensional profiles (section 25.12), but such documents are not yet publicly available for three-dimensional areal surfaces. Surface texture parameters can be broadly classified into two main groups; field parameters that utilise the measured data and feature parameters that use a subset of it. Feature parameters describe the characteristics of pits, peaks and saddle points. Quantification of surface texture is especially critical for certain engineering applications, for example, in designing the lubricated surface of a piston moving within a cylinder. This need has led to a subset of parameters being developed specifically for engineers, e.g., the core fluid retention index. Although these quantities appear to have little relevance for biomaterials they may be important for characterising microcracks that could act as anchor points for pseudopodia or be relevant for protein adsorption studies.

25.7 Two-dimensional profile data Parameters that provide useful descriptions of surface profiles are grouped as: · amplitude parameters that are measures of variations in profile height · spatial parameters that describe in-plane variations in surface texture · hybrid parameters that combine both amplitude and spatial information, e.g., mean slope.

25.7.1 Amplitude parameters Amplitude parameters can be grouped into two subclasses: averaging parameters and peak to valley parameters.

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25.2 (a) Averaging the peaks and troughs in measured profile data over a given length scale is used to generate a midline. (b) The grey valleys are inverted to form peaks that are averaged with the existing peaks to obtain Ra, the arithmetic mean deviation from the midline.

Averaging parameters The most widely used parameter to quantify surface texture is Ra the arithmetical mean deviation of the absolute ordinate values, z…x† of the profile from a midline: Z 1 lr jz…x†j dx 25:1 Ra ˆ lr 0 where lr is the sampling length over which the surface profile has been measured. The derivation of Ra is schematically shown in Fig. 25.2. The Ra parameter is meaningless if given without reference to the cut-off filter wavelength, c (section 25.5). Despite its popularity the Ra value does not provide detailed information about the geometry of the surface or the variations in peak heights or valley depths. Because of this lack of information Ra is not a useful parameter to use for assessing the texture of biomaterial surfaces, a view that is endorsed by Fig. 25.3, which shows how different surface profiles can have equivalent Ra values. The parameter, Ra is typically used in engineering to describe the roughness of machined surfaces and is a useful quality control parameter for monitoring established manufacturing processes. The statistical significance of Ra is improved by averaging the values obtained for each of the five sampling lengths that comprise the evaluation length. If Ra is determined

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25.3 A schematic representation of how quite different surface textures (a, b and c) can have the same Ra value.

from other than five sampling lengths then it is written as RaX, where X is the number of sampling lengths used in the averaging process. Some of the limitations of the arithmetical mean deviation of the profile, Ra, can be overcome by using Rq, the root-mean-square value of all distances of the measured profile away from the centre line. Rq is given by: s Z 1 lr 2 Rq ˆ z …x† dx 25:2 lr 0 Conceptually Rq is very similar to Ra, however, the critical difference between the two parameters is that the deviations of the peak heights and valley depths from the midline appear as a squared term in Rq. This increases the sensitivity, or susceptibility, of this parameter to outlying points, i.e., dominant peaks and/or valleys. Peak and valley parameters The distribution of peak heights and valley depths can provide valuable information about surface texture. A surface that has both a wide range of peak heights and valley depths will have a probabilistic bell-shaped distribution centred on the mean. The dimensionless skewness parameter, Rsk, is used to detect and quantify bias in the shape of this distribution where: Z lr 1 z…x†3 dx 25:3 Rsk ˆ 3 Rq lr 0 The skewness of a perfectly random surface with a wide range of peak heights and valley depths is zero. Wear of such a surface will tend to smooth out the peaks whilst leaving the valleys intact skewing the distribution away from zero to negative values. The converse is true for processes such as electroplating that

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25.4 The relationship between the maximum peak height and valley depth within a profile.

tend to fill the valleys but leave the peaks intact. Porous materials that have a preponderance of `valleys' have negative values of Rsk. Changes in this parameter could, for example, be used to monitor time-dependent biodegradation of materials. Kurtosis, Rku, is a statistical measure of the sharpness of a histogram of peak heights and valley depths: Z lr 1 Rku ˆ 4 z…x†4 dx 25:4 Rq lr 0 More specifically Rku is a measure of the comparison of the measured profile with the Gaussian distribution characteristic of a perfectly random distribution of peak heights and valley depths. The kurtosis of a Gaussian distribution is three. Sharper profiles have higher values of kurtosis and broader, flatter profiles have kurtosis values that are less than three. Figure 25.4 shows the relationship between some commonly used amplitude parameters and surface texture; Rp is the maximum profile height above the mean line in the measured profile and Rv is the minimum profile depth below the mean line. Rz is the distance between the two. Clearly Rp and Rv need to be identified very carefully as surface damage, the presence of foreign bodies and sampling position, can all lead to a misleading representation of surface texture.

25.7.2 Spacing and hybrid parameters Some spatial information can be obtained through the mean width of the profile elements, RSm. RSm is the mean value of the profile element widths, Xs shown in Fig. 25.5, i.e., the average width of a peak and valley combined. The position

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25.5 The mean width of the profile elements Xsi is RSm.

of the line along which these data are derived from is stated in ISO 4287: 1997 as being 10% of Rz with a default sampling of 1% of the sampling length. Techniques such as the quartz crystal microbalance are used to detect, for example, protein deposition onto surfaces through changes in mass and hence the resonant frequency of the quartz crystal. It is generally assumed that the protein coating is of single molecule thickness and, therefore, any differences in the mass of protein adsorbed onto different surfaces is due to differences in its surface affinity. Whilst this may be true it does assume that the area available for protein adsorption is constant for each surface studied. The hybrid parameter, R
q can be used to find the actual length of a profile and hence be used to compare the actual length of surface available for molecular adsorption. R
q is the root mean square slope, dZ=dX , over the length of the profile at a location Y on the surface. It is defined as a hybrid parameter since it contains both amplitude and spacing information.

25.7.3 Summary of two-dimensional parameters The two-dimensional parameters that can be used to characterise the surface texture of biomaterial surfaces are listed in Table 25.1.

25.8 Three-dimensional data A limitation of two-dimensional profiles is that the information produced may not be representative of the whole sample. This uncertainty can be reduced by either averaging a number of profiles or by directly analysing a region of the sample in three dimensions. The three-dimensional areal parameters were developed during the 1980s and led to a proliferation of parameters being proposed that Whitehouse described as `parameter rash' (Whitehouse, 1982). Since that time a number of

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Table 25.1 Two-dimensional parameters for quantifying the texture of biomaterial surfaces (derived from Blunt 2001) Ra Rq Rsk Rku

Amplitude parameters Arithmetical mean deviation of the profile The root-mean square (RMS) of the deviation of the profile (m) Skewness of the profile Kurtosis of the profile

Rp Rv Rz

Specific peak and valley parameters Maximum peak height Maximum valley depth Distance between Rp and Rv

RSm R
q

Spacing and hybrid parameters Mean width of profile elements RMS slope of the assessed profile

initiatives (Stout et al., 1993) have led to a rationalisation of the number of parameters so that they are closely related to those used to quantify twodimensional surface profiles (Blunt, 2001). Specification standards that define these areal parameters and their applications are currently being developed within ISO (the International Organisation for Standards). Areal parameters that describe surface texture are based on their twodimensional equivalents and are similarly broken down into amplitude, spatial and hybrid parameters; another class has been described but this has little relevance to biomaterials. The three-dimensional parameters are particularly valuable for identifying patterns in surface texture indicated by high values of Str (the texture aspect ratio), and low values for Sal (the fastest decay autocorrelation length).

25.8.1 Amplitude parameters Many of the amplitude parameters used to describe three-dimensional surface texture are simply derived from their two-dimensional equivalents. Parameters such as the maximum peak height and minimum valley depth are not particularly valuable parameters because they have a tendency to isolate surface features such as dirt and scratches. Even the three-dimensional equivalent of Ra is considered to be less useful than the statistically more significant root-meansquared departure from a midline, Sq. The Sq parameter is given by: s ZZ 1 z2 …x; y† dxdy 25:5 Sq ˆ A A where A is the measurement area. The distribution of surface heights amplitudes can be assessed using measures of skewness, Ssk and kurtosis, Sku following the analysis of two-dimensional data. Three-dimensional skewness is given by:

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25:6

and kurtosis by:

 ZZ  1 1 4 z …x; y† dxdy : Sku ˆ 4 Sq A A

25:7

25.8.2 Texture and hybrid parameters The density of summits, Sds, is a measure of the number of peaks within a unit sampling area Sds ˆ

Number of peaks A

25:8

The Sds parameter is useful for assessing applications that involve friction and specific, or close but non-contacting proximity, of surfaces. In traditional engineering this means the assessment of bearing surfaces for signs of wear, which when transferred to biomaterials can refer to the wear of articulating surfaces, e.g., a hip joint due to fretting (Howell et al., 2000). The texture aspect ratio of the surface, Str, is used to identify how consistently uniform the surface texture is. Str is given by a complex ratio of autocorrelation lengths that is beyond the scope of this chapter. In practice the texture aspect ratio varies from 0 to 1. Values of Str in excess of 0.5 indicate that the surface texture is uniform in all directions, i.e., the texture has no defined lay, unlike the baize of a snooker table. Values of Str below 0.3 imply increasing amounts of directional structure in the texture, i.e., the development of knap or lay. This parameter for characterising biomaterial surfaces has potential relevance in identifying the remnants of machining or tooling marks that may influence cell or protein adhesion. The texture direction of the surface, Std, indicates the direction of the dominant lay with respect to the y-axis. When the lay direction is perpendicular to the measurement trace direction, Std is 0ë. The fastest decay autocorrelation length, Sal, is used to describe the behaviour of the autocorrelation function. More specifically Sal is a measure of the distance over which the autocorrelation function decays to its value of roughly 0.2. Large values of Sal are characteristic of surfaces that are dominated by long wavelength spatial features, the opposite being true for small values of Sal. The autocorrelation function describes the occurrence of regular undulations of the surface revealing dominant wavelengths that lie within the surface texture. Hybrid parameters contain elements of both amplitude and spatial information. The root-mean-square slope of the surface, Sdq, for example, is a measure of the surface slope over the sampled area.

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Table 25.2 Three-dimensional parameters for quantifying surface texture (based on Blunt, 2001)

Ssk Sku

Amplitude parameters The root-mean square of the deviation of the surface (m) Sum of the largest peak and deepest pit within a defined area (m) Skewness of the surface Kurtosis of the surface

Str Sal

Spatial parameters Texture aspect ratio Fastest decay autocorrelation length (mm)

Sdq Sds Sdr

Hybrid parameters Root-mean square slope of the surface (m/m) Density of summits per unit area Developed surface area ratio (%)

Sq Sz

25.8.3 Summary of three-dimensional parameters The three-dimensional parameters used to characterise surface texture are listed in Table 25.2.

25.9 Techniques for surface texture measurement A number of techniques are available for measuring surface texture that include both contact and non-contacting methods. The principles of these different approaches are described below. It is important to appreciate that each of these methods has its limitations in terms of what can and cannot be measured, e.g., optical instruments are unable to detect undercuts, stylus instruments are unable to precisely trace sharp peaks or to penetrate steep-sided valleys. In addition to these physical limitations there are also instrumentation issues, i.e., the damping characteristics of the cantilevers used in stylus instruments, the wavelength of light used in optical profilers, all of which limit the accuracy of surface topographical measurements.

25.9.1 Contact measurements Stylus instruments The basic characteristics of a contact (stylus) instrument are described in ISO 3274: 1996. The principle is analogous to that of a gramophone stylus moving across a record; movement of the stylus over a textured surface induces movement of the stylus that can be measured and used to derive positional information (Fig. 25.6). Movement of the stylus arm can be detected using a

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25.6 Schematic representation of a stylus contact device.

piezoelectric crystal or moving coil device. These devices are suitable only for dynamic measurements, as these types of detectors generate a voltage in response to movement of the armature. Measurements of the static stylus position can be obtained from a linear variable differential transformer (LVDT) by measuring the inductive current generated in an AC energised static coil by the proximity of one mounted on the armature. This detection system can be used for both static and dynamic movements of the stylus. Stylus movement can also be detected by coupling the armature to a laser interferometer. In this arrangement a reflector mounted on the end of the armature changes the path length of a split beam of light that, when combined with that reflected from a fixed mirror, produces optical interference. Typically cone-shaped styli with spherical tips are used for measuring surface features. The cone angle is usually 60ë or 90ë with tip radii ranging from 1 m to 10 m. Both the cone angle and tip radius will affect the sensitivity of the stylus, Table 25.3. Stylus instruments tend to distort surface features by rounding-off peaks and reducing the depth of valleys within the surface as schematically shown in Fig. 25.7. Contact stylus instruments have other limitations; they are unable to · follow overhang features (Fig. 25.8) · penetrate steep-sided valleys; this becomes increasingly important as the tip radius increases in size with respect to the valley dimensions · measure the surface features of soft materials without damaging the surface. Table 25.3 The relationship between the low-pass filter cut-off wavelength, c, the high-pass filter cut-off, s, the target radius of the probe tip (Rtip max) and the sampling interval (from ISO 3274: 1996) c (m)

s (m)

Roughness cut-off wavelength ratio c/s

Rtip max (m)

Maximum sampling spacing (m)

0.08 0.25 0.80 2.50 8.00

2.5 2.5 2.5 8.0 25.0

30 100 300 300 300

2 2 2 5 10

0.5 0.5 0.5 1.5 5.0

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25.7 Stylus instruments tend to round-off peaks and reduce the depth of valleys.

25.8 Stylus instruments are unable to accommodate overhangs.

Stylus instruments are, however, relatively inexpensive and straightforward to use and can scan strips up to hundreds of millimetres in length. Used with laser interferometers to detect movement of the armature, stylus instruments have an out-of-plane resolution of around 10 nm or better over a sampling length of 6 mm. Currently the stylus approach to measuring surface texture is the only surface texture measurement method covered by international standards. ISO 3274: 1996 recommends a stylus force of 0.75 mN to avoid damaging the sample although this is difficult to set in practice.

25.9.2 Non-contacting measurements Optical methods Optical profilers are especially useful for assessing the surface roughness of soft materials, as there is no physical contact between the probe and the surface. They can provide much shorter measurement times, especially in three dimensions, than mechanical devices. Optical techniques can be categorised

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into two groups depending on the spatial coherence of the light source. Monochromatic sources, e.g., lasers are typically used in constant focus devices (profilometers), phase shifting interferometry, holography, fringe projection and speckle techniques. Polychromatic light sources are also used in white light interferometers, interference microscopy and in coherence scanning methods. There are practical limitations in using optical methods; one of the most important being that the surface needs to be reflective. The minimum reflectivity can be as low as ~4% but at these levels weak signals can lead to `holes' in the interferometric data and disorientation of focusing devices. The reflectivity of matt surfaces can be significantly improved by applying a fine coat of paint, lycopodium powder or by sputter coating a layer of gold over the surface. Coating needs to be done with caution to avoid significantly altering the surface features; a problem that becomes increasingly significant with increasing sensitivity of the measurement. Slopes that are steeper than around 15ë are difficult to measure although many of the software packages provided with the measuring instrument allow the user to interpolate data to `patch' holes. The level of sophistication associated with this process is both software and user dependent. Materials that are composed of more than one component can also be difficult to measure using optical techniques, especially if there are significant differences in the optical constants. Optical focusing techniques Laser profilometers rely on monitoring the movement of an objective lens as it moves across the specimen surfaces in the z (out-of-plane) direction to maintain constant focus. The beam of a laser profilometer is typically 1 m to 2 m in diameter with a vertical resolution of around 100 m, although some products are available that claim to have much smaller resolution limits. Profilometers have difficulty in measuring surfaces that have steep steps or discontinuities due to an irrecoverable loss of focus. White light interferometry Figure 25.9 illustrates a typical white light interferometer (WLI) configuration. The vertical scanning WLI uses a broadband light source and measures the degree of modulation contrast as a function of path difference. Because of the large spectral bandwidth of the source, the temporal coherence length of the source is short, so high contrast fringes will be obtained only when the two paths in the interferometer are closely matched in length. So by modifying and measuring the path difference while monitoring the fringe contrast and phase of the interference for each pixel on a CCD detector, the height variations across a surface can be determined.

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25.9 A schematic representation of a white light interferometer.

Vertical scanning white light interferometers usually have multiple modes of operation depending on the surface being measured. Vertical resolutions are less than 1 nm and the horizontal resolution is determined ultimately by the wavelength of the source. The vertical range is usually around 100 m and the horizontal range will be determined by the objective lens used. However, both vertical and horizontal ranges can be extended to millimetres of range by using mechanical scanning stages and digital image stitching techniques (with corresponding losses in accuracy). Despite the relatively simple theory behind the operation of the vertical scanning WLI, there are problems that can occur in practice that are not always obvious to the user. Some examples include the effects of different materials present on the surface (phase change effects) and the sensitivity to slopes at the surface due to the finite numerical aperture of the interferometer and diffraction. Vertical scanning WLI interferometers are very versatile and can be used to measure the surfaces of materials ranging from low reflectivity plastics to high reflectivity mirrors. The main advantage of the WLI over the stylus instrument is the speed of the measurement. Typically a WLI will take less than a minute to take a measurement that may take hours on a stylus instrument.

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Electronic speckle pattern interferometry (ESPI) An optically rough surface when illuminated by a coherent light source, i.e., a laser, appears speckled. These speckles can be used to investigate the topography of the surface using electronic speckle pattern interferometry (ESPI) (Jones and Wykes, 1989). This technique was `invented' some 30 years ago and is still finding new applications; recently the principle has been adapted to simultaneously measure both in- and out-of-plane displacements (Hurtado-Ramos et al., 2001). The principle of the technique is to illuminate a surface using a defocused laser beam (Fig. 25.10). Reflections from the surface are combined with a reference beam to produce a speckle pattern that is recorded by a video camera. The sample is then displaced and a second speckle interferogram is recorded. The difference between the two interference patterns is related to movement of the specimen at any given position on its surface. For static measurements of roughness illuminating the specimen sequentially using two different wavelengths that have a wavelength difference of around 10 nm generates the two interferograms. A interference pattern is obtained by subtracting the two interferograms. This pattern can then be analysed to produce a displacement map. An alternative approach to using two wavelengths is to move the position of the light source or the sample to generate the interference pattern. This technique, although fast, is suitable only for continuous surfaces where surface features cause phase shifts of less than 2. In more practical terms, this technique is not able to detect features that are larger than the wavelength of the light source between two adjacent points.

25.10 A surface illuminated by two de-focused lasers produces a speckle pattern.

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Scanning probe microscopy Scanning probe microscopy covers several techniques for imaging surfaces down to the atomic level as well as scanning on a much broader scale, e.g., 100 m by 100 m, assuming that the height variations are below approximately 10 m. All of these techniques rely on monitoring the interaction of a fine tipped probe, typically having a radius of curvature lying between 3 nm and 50 nm, with the surface. Commonly used techniques are: · atomic force microscopy, where the interactive force between the tip and surface is monitored as the probe passes over the surface · scanning tunnelling microscopy, where the flow of an electrical tunnelling current is monitored between the tip and the surface as it glides over the surface at constant height · near-field scanning optical microscopy, where the sample is illuminated by light from a laser beam that has passed through a sub-wavelength hole. This arrangement extends the capability of light microscopes to image length scales below 50 nm. Atomic force microscopy (AFM) Atomic force microscopy (Fig. 25.11) is used to image biomaterials at the atomic level. Such detailed information is particularly relevant for studies of

25.11 In atomic force microscopy measurements of the interactive force between an atomic resolution tip and a surface give rise to an image.

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Table 25.4 The atomic microscope can be used in different modes to measure different properties of the surface. These methods are commonly used to investigate biological materials Mode

Nature of the probe/ surface interaction

Comments

Contact ± probe in contact with surface

Strong repulsive force

Probe maintained at either constant distance or positioned to give a constant force

Non-contact

Weak, attractive forces

Probe vibrates

Intermittent contact (tapping modeTM)

Strong repulsive force

Probe vibrates

Phase

Used to measure the phase difference between different measured modes, e.g., frictional force and contact profile.

protein adsorption or cell adhesion to subtrates. Atomic force microscopy is used in several modes that differ according to the force between the probe tip and the surface (Table 25.4). In contact mode the AFM behaves like a miniature stylus contact instrument (section 25.9.1). Like any stylus instrument in contact mode the AFM tip is not well suited to characterising the surface texture of soft materials. Surface debris can also pose significant problems leading to tip damage or blunting. These problems can be overcome by using a near-contact mode. In this mode the AFM is particularly susceptible to vibration and to any moisture present on the surface as an adsorbed layer; a common problem for biomaterials in uncontrolled atmospheres. The moisture film through capillarity tends to resist movement of the cantilever as it is pulled away from the surface leading to measurement problems. Despite these limitations near-contact modes are particularly well suited for imaging soft materials. Any external vibration including that attributed to passing traffic and people will affect the performance of the atomic force microscope. The sensitivity of the device also depends on the size and shape of the probe tip and on the quality of the mechanical positioning elements that are used to control the position of the probe. Care must be taken to select the piezoelectric drive unit that is usually the heart of the positioning unit so that it operates linearly over the usable range and does not overshoot during rapid movement. A number of nano-fabricated artefacts have been developed for calibrating AFMs that are typically gratings with well-defined step heights.

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At high magnifications the shape of the tip is important. Sharp tips improve the resolution of the instrument since interactive forces between the tip and the underlying substrate decay away relatively slowly with increasing radial distance away from the point of the tip. This has an averaging effect with near neighbour atoms influencing the measured force. The effect becomes increasingly important as the probe tip becomes less sharp. The AFM produces an image that is effectively an `averaged out' surface; the significance of this for someone investigating biomaterials will depend on the intended application. Even very sharp tips can cause smearing in the out-of-plane direction for abrupt changes in height akin to that produced by conventional stylus probe instruments leading to `false' images. Lateral forces between the probe tip and the sample present in contact mode AFM pose the major problem in imaging soft matter. This problem is virtually eliminated in techniques in which the cantilever is oscillated above the surface (ac techniques), and the topography is tracked by monitoring the cantilever oscillation (amplitude, frequency, phase) under the influence of tip-sample interactions. The set-up in which the non-contact interaction is monitored might appear to be most desirable from the point of view of maintaining the noninvasive character of operation. However, the intermittent-contact technique, in which the tip taps on the sample surface at the bottom of each oscillation cycle (tapping mode), is currently most widely used. The resolution of the AFM is typically around 0.1 nm in the vertical out-ofplane direction and several nanometres in the in-plane axis. The in-plane range of the AFM is approximately 100 m by 100 m. Scanning tunnelling microscopy (STM) Scanning tunnelling microscopy is used to produce very high-resolution images of surface atoms. These devices function by positioning a sharp probe tip, usually made of metal and theoretically terminating with a single atom, about 1 nm from the surface being measured. At this very small separation a tunnelling current flows between the tip and the substrate which is maintained at a potential difference of between 1 mV to 3 mV. The tunnelling current is very small ranging from picoamperes to nanoamperes and flows across the gap between the surface and the tip. As the tip moves over the surface a piezoelectric controller maintains a constant separation so that the tunnelling current is also constant. Movement of the tip in the out-of-plane axis is monitored via the voltage applied to the piezoelectric controller. This is used to generate an image of the surface topography. Ideally this method should only map the variations in surface topography, but in practice the image also reflects the variation in electronic density of states across the surface of the sample and in interactions between the surface and the tip that have yet to be understood. STM can also be used in constant height mode where the probe height above the surface is maintained and the variation in current flow is used to map the topography. This approach,

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whilst good for smooth surfaces, is risky to use for rough surfaces. STM can be used in air, liquids and under vacuum which enhances its value in investigating biomaterial surfaces. Surface contamination during scans can be problematical hence most measurements are made under ultrahigh vacuum. STM has a vertical resolution in the picometre range and a lateral resolution of approximately 0.1 nm. The vertical range of the STM is around 5 m and the inplane range is approximately 100 m by 100 m. A drawback of STM is that it can only be used to measure the topography of conducting surfaces. There is also some ambiguity in the origin of the tunnelling current. The user relies on the fact that the tunnelling current involves a single atom, but this need not be the case. If the probe tip passes over a slope or pit then the tunnelling point can change to another atom near the slope surface or pit giving rise to an erroneous reading. Scanning near-field optical microscopy According to Abbe, the resolving power of a classical light microscope is limited to half the wavelength of the light source, i.e., around 200 nm, however, new light microscopes have been built that are able to overcome this limitation. In the scanning near-field microscope (SNOM) light from a laser beam passes through a sub-wavelength diameter aperture to illuminate a specimen placed within its near field at a distance that is shorter than the wavelength of light. The resolving power of such a microscope is less than 50 nm. The light source needs to be brought to within a few nanometres of the surface in order for the technique to work and the reflected light is then collected and detected. In operation the light source moves over the sample surface without touching it. The relevance of this technique for assessing the surface texture of biomaterials is more orientated towards studies of protein adhesion as the surface being assessed needs to be smooth at the nanometre level. Several methods are used to generate the point source of light including: · focusing a laser beam through a sub-wavelength diameter hole drilled through the centre of an AFM tip · using metal coated tapered optical fibres that terminate in an uncoated tip 50 nm or less in diameter · filling the tip of a very finely tapered pipette with a compound that emits light under an applied voltage. The size of the aperture in each of the point sources corresponds to the lateral resolution of the microscope. Four modes of operation are used with SNOM (Fig. 25.12): · Transmission. A photomultiplier tube or similar device positioned beneath the sample collects and detects light that passes through it.

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25.12 Modes of operation for SNOM.

· Reflection. The sample is locally illuminated by light passing through the probe and the reflected light is collected and detected. · Collection. A light source is used to illuminate the sample from beneath. The SNOM probe is used to collect and detect light that has been transmitted through the sample. · Illumination/collection. The probe is used both to deliver light to the sample surface and to collect and detect that reflected from it. The sample image is usually derived from changes in the intensity of the reflected or transmitted light that occurs as the point light source moves across the surface of the sample. Changes in other optical properties such as polarisation, refractive index or the presence of fluorescent markers can also be used to generate images.

25.10 Traceability and calibration It is good practice to ensure that any measurements of surface texture are made using instruments that are regularly calibrated using procedures that can be traced back to national or international standards. This process ensures that the data are accurate to within a calculable uncertainty (Leach, 2001).

25.11 Conclusion A variety of techniques can be used to measure the surface features of materials. These can be grouped into non-contacting, e.g., imaging methods or contacting techniques, e.g., stylus probes. Some of these methods are complementary; all have their limitations. The electron microscope can be used to image surfaces to visually assess roughness, however, the standard technique is unable to generate quantitative data in the out-of-plane direction, therefore, it is of limited value in comparing different surface textures. Techniques are available for measuring the coordinates in two or three dimensions from rough surfaces that include the out-of-plane dimension. These include stylus probes and non-contacting methods such as white light interferometry and laser profilometry. Stylus methods are particularly well suited to

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measuring the surface roughness of hard materials such as ceramics or metals but can cause `ploughing furrows' in softer materials. The non-contacting approach is more appropriate for plastics although caution is required to ensure that the data generated is reliable. Laser profilometers are used to measure relatively coarse surfaces with resolution limits of around 1 m, a value that depends on the sampling frequency, the filtration used and the sensitivity of the transducer. Laser profilometers are particularly valuable for measuring the form of a surface, i.e., the broad texture. Finer detail, that may be important for cell adhesion, as well as form information, can be obtained from a white light interferometer where the resolution limit is in the nanometre range, a factor that depends on the flatness of the sample and the objective used. Both the white light interferometer and laser profilometer have limitations in the angular changes that they can detect, sharp edges cause problems for both techniques. The cut-off point for detection depends on the numerical aperture of the objective used for both instruments.

25.12 Further reading Three dimensional surface topography, K.J. Stout and L. Blunt, Penton Press, 2000. Rough Surfaces, 2nd edn, T.R., Thomas, Imperial College Press, London, 1999.

25.12.1 Scanning probe microscopy Atomic Force Microscopy for Biologists; V.J. Morris, A.R. Kirby, A.P. Gunning, Imperial College Press; ISBN: 1860941990 (December 1999). Atomic Force Microscopy: Biomedical Methods and Applications, P.C. Braga, D. Ricci, Humana Pr; ISBN: 1588290948, (1 October, 2003). Scanning Probe Microscopes: Applications in Science and Technology, K.S. Birdi, CRC Press; ISBN: 0849309301, (31 August, 2002). Scanning Probe Microscopy of Polymers (ACS Symposium Series), B.D. Ratner and V.V. Tsukruk (eds), American Chemical Society; ISBN: 0841235627, (27 August, 1998). The Measurement of Surface Texture using Stylus Instruments, R. Leach, National Physical Laboratory Measurement Good Practice Guide No. 37. 2001. (This document contains a list both of instrument distributors and UKAS accredited laboratories for surface texture measurement.)

25.12.2 Standards ISO 1302: 2002 Geometrical Product Specifications (GPS) ± indication of surface texture in technical product documentation. ISO 3274: 1996 Geometrical Product Specifications (GPS) ± surface texture: profile method ± nominal characteristics of contact (stylus) instruments. ISO 4287: 1997 Geometrical Product Specifications (GPS) ± surface texture: profile method ± terms, definitions and surface texture parameters. ISO 4288: 1996 Geometrical Product Specifications (GPS) ± surface texture: profile method ± rules and procedures for the assessment of surface texture.

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ISO 5436-1: 2000 Geometrical Product Specifications (GPS) ± surface texture: profile method; measurement standards ± Part 1: Material measures. ISO 5436-2: 2001 Geometrical Product Specifications (GPS) ± surface texture: profile method; measurement standards ± Part 2: Software measurement standards. ISO 11562: 1996 Geometrical Product Specifications (GPS) ± Surface texture: profile method ± metrological characteristics of phase correct filters. ISO 12085: 1996 Geometrical Product Specifications (GPS) ± surface texture: profile method ± motif parameters. ISO 12179: 2000 Geometrical Product Specifications (GPS) ± surface texture: ISO 13565-1: 1996 Geometrical Product Specifications (GPS) ± surface texture: profile method; surfaces having stratified functional properties ± Part 1: Filtering and general measurement conditions. ISO 13565-2: 1996 Geometrical Product Specifications (GPS) ± surface texture: profile method; surfaces having stratified functional properties ± Part 2: Height characterisation using the linear material ratio curve. ISO 13565-3: 1998 Geometrical Product Specifications (GPS) ± surface texture: profile method; surfaces having stratified functional properties ± Part 3: Height characterisation using the material probability curve.

25.13 Acknowledgements This work was funded by the United Kingdom Department of Trade and Industry as part of its programme of research on Materials for Processing and Performance (Project MPP 4.5).

25.14 References Blunt, L., The final report: The development of a basis for 3D surface roughness standards, (EC SMT$-CT98-2256), July 2001. Bowers, K., Keller, J.C., Randolph, B., Wick, D. and Michaels, C., `Optimization of surface micromorphology for enhanced osteoblast responses in vitro'. Int. J. Oral Maxillofac. Impl. 7, 302±310, 1992. Hallab, N.J., Bundy, K.J., O'Connor, K., Moses, R.L., and Jacobs, J.J., `Evaluation of metallic and polymeric biomaterial surface energy and surface roughness characteristics for directed cell adhesion', Tissue Engineering, 7(1), 55, 2001. Howell, J.R., Blunt, L.A., Lee, A.J.C., Hooper, R.M., Gie, G.A., Timperley, A.J. and Ling, R.S.M., `An investigation of the fretting wear seen on explanted hip replacement femoral stems', J. Bone Surg., 82(B), 52±53, 2000. Hurtado-Ramos, J.B., Blanco-Garcia, J., Fernandez, A. and Ribas, F., `An ESPI system for determining in-plane deformations. Three-dimensional analysis of the carrier fringes and a proposal for the analysis of transient in-plane deformations', Meas. Sci. Technol, 12, 644±651, 2001. Jones, R. and Wykes, C., Holographic and Speckle Interferometry, 2nd edn. Cambridge University Press, 1989. Keller, J., Stanford, C.M., Wightman, J.P., Draughn, R.A. and Zaharias, R., `Characterisation of titanium implant surfaces 111', J. Biomed. Mater. Res. 28, 939±946, 1994.

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Leach, R., The Measurement of Surface Texture using Stylus Instruments, National Physical Laboratory Measurement Good Practice Guide No. 37. 2001. Ogawa, T., Sukotjo, C. and Nishimura, I., `Modulated bone matrix-related gene expression is associated with differences in interfacial strength of different implant surface roughness', J Prosthodont. 2002 Dec; 11(4): 241±7. Rosa, A.L. and Beloti, M.M., `Effect of cpTi surface roughness on human bone marrow cell attachment, proliferation and differentiation', Braz. Dent. J. 14(1), 16±21, 2003. Schneider, G.B., Perinpanayagam, H., Clegg, M., Zaharias, R., Seabold, D., Keller, J. and Stanford, C., `Implant surface roughness affects osteoblast gene expression', J Dent Res., 2003, 82(5), 372±6, May. Stout, K.J., Sullivan, P.J., Dong, W.P., Mainsah, N., Luo, N. Mathia, T. and Zahyouani, H., The development of methods for the characterisation of roughness in three dimensions, 1st edn Commission of the European Communities (ISBN 0 70441 3131 2), 1993. Whitehouse, D.J., `The parameter rash ± is there a cure?' Wear, 83, 75±78, 1982.

Part V Appendices

26

Surface modification of polymers to enhance biocompatibility M T A V A K O L I , TWI Limited, UK

26.1 Introduction In spite of the availability of many polymers as medical materials, the biocompatibility of these materials particularly in long-term applications is limited. This has generated considerable interest in the research community to produce polymers with enhanced biocompatibility. When a biomaterial contacts a cell, tissue, blood or other biological fluid, it is the surface of the material that comes into contact with the physiological environment. Many medical materials and biomaterials, semi- or permanent implantable medical devices, are made from polymers and the biocompatibility of these materials is critical to their satisfactory performance in the intended application. In general, the first physiological process that occurs within the initial stages of exposure is the adsorption of biomolecules onto the surface and this may be followed by cellular interactions, if cells are encountered. This has led to considerable effort to modify the surfaces of medical materials to enhance their biocompatibility. The main activities have focused on the development of novel surface modification techniques for the adhesion of various cells, enhancement of tissue biocompatibility as well as haemocompatibility enhancement of polymer surfaces. A variety of different techniques have been explored to modify polymer surfaces in order to induce specific surface property and bio- and haemocompatibility. These include: · · · · · ·

plasma lasers photolithography surface grafting coating other techniques for haemocompatibility.

These techniques will allow the design and preparation of new polymers for medical and biomaterial applications with specific surface properties such as:

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biocompatibility and haemocompatibility wear resistance permeability/impermeability electrical insulation/conducting properties.

A review of some of the most promising surface modification techniques will be given in this chapter.

26.2 Polymers in medical applications Naturally occurring and synthetic polymers have been used or considered for many applications as medical materials and biomaterials.1±4 Compared to other materials, the use of polymers in medical applications has increased significantly in recent years. This is primarily based on their wide range of chemical, physical and mechanical properties, their processability by a variety of different techniques and their availability in different forms and shapes. The unique properties of polymers enable them to undergo controlled biodegradation or to be bioresorbable or bioresistant, offering significant advantages over their metallic and ceramic counterparts. Polymers as medical materials and biomaterials are being used in many applications from dentistry (acrylic based dental materials) to pharmacy (e.g. cellulose and acrylics in drug delivery systems) and surgery (e.g. cyanoacrylate based tissue adhesives) and as scaffolding in tissue engineering. However, the major use of polymers is in the manufacture of a variety of disposable, short- and long-term implantable medical devices.

26.2.1 Classification Polymeric materials can be mainly divided into three major classes: thermosets, thermoplastics and elastomers (Fig. 26.1). A simple illustration of the definition of thermoplastic and thermosetting polymers is shown in Fig. 26.2. A thermosetting polymer, as the name suggests, becomes set into a given network, normally through the action of a catalyst, heat, radiation or combination of these

26.1 General classification of polymers.

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26.2 An illustration of the difference between thermoplastic and thermosetting polymers.

factors, via the process of crosslinking. As a result of this, thermosets become infusible and insoluble. Thermosetting resins (e.g. epoxies) are the basis of many medical grade adhesives for joining materials in a number of medical devices. In contrast, thermoplastic polymers (e.g. polyethylene (PE)) may be defined as materials, which soften, melt and flow on the application of heat and solidify on cooling. Thermoplastics (e.g. PE, polyurethanes (PU), polyvinylchloride (PVC), PE, polyetheretherketone (PEEK)) are the basis of many medical polymers used in disposable and implantable devices. An illustration of typical applications of polymers in medical applications is shown in Fig. 26.3.

26.3 Typical examples of uses of polymers in medical applications.

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26.2.2 Compliance with USP and ISO Standards Tests to determine the biological reactivity of polymeric materials and medical devices are described in the USP (United States Pharmacopeia) and ISO (EN ISO 10993-1) standards.5 According to the injection and implantation testing requirements specified under Biological Reactivity Tests, in-vivo polymers are scaled on a Class of I to VI. To grade a polymer, extracts of the test material are generated in various media and injected systematically and intracutaneously into rabbits or mice to evaluate their biocompatibility. Polymers not requiring implantation are graded Class I, II, III or V and those polymers requiring implantation testing are usually graded Class IV or VI. Many polymers are available which can be qualified as USP Class IV and VI materials. These materials can pass the relevant intracutaneous toxicity (in vivo), acute systemic toxicity (in vivo) and implantation (in vivo) testing requirements. Passing USP Class VI standards does not guarantee that a polymer will meet the FDA requirements in a particular application. However, passing the test is a strong indication of the non-toxicity of a polymer. The ISO Standard 10993 currently consists of 18 parts (17 approved and Part 18 under consideration for approval). Part one describes guidance on selection of tests. Part two includes animal welfare requirements. Parts 3±18 describe specific tests, which include a variety of toxicity and other tests for identification and quantification of degradation products from polymers. Many medical grade polymers are capable of withstanding exposure to less severe sterilisation procedures (e.g. gamma irradiation and ethylene oxide). However, other types of thermoplastic (e.g. PEEK) or thermosetting (e.g. polyimides) polymers can be exposed to more sever sterilisation (e.g. autoclaving) as well as gamma irradiation, ethylene oxide (ETO) and chemical methods.

26.3 Biocompatibility As Williams, the originator of the definition of biocompatibility describes,6 this is the ability of a material to perform with an appropriate response in a specific application. This definition separates biocompatibility from biological safety. In a more recent publication,7 Williams suggests that it would be better to base the definition of biocompatibility on the application of a device and not on the material. Williams suggested the following as the definitions of short-term and long-term implantable devices:7,8 Short-term: `The biocompatibility of a medical device that is intentionally placed within the cardiovascular system for transient diagnostic or therapeutic purposes, refers to the ability of the device to carry out its intended function within flowing blood, with minimal interaction between device and blood that adversely affects device performance, and without including uncontrolled activation of cellular or plasma protein cascades.'

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Long-term: `The biocompatibility of a long-term implantable device refers to the ability of the device to perform its intended function, with desired degree of incorporation in the host, without eliciting any undesirable local or systemic effects in that host.'

A medical polymer could satisfy many biological safety tests but fail biocompatibility requirements and so be considered to be bio-incompatible in some applications. There are many organic (e.g. polymers) and inorganic materials (e.g. titanium, ceramics) that are available, which are qualified for applications for short-term (<30 days) or long-term (>30 days) implantation. These materials may already have a certificate of compliance to FDA or European Standards. For instance, for short-term (<30 days) contact or implantation inside the body there are USP and ISO standards. Polymeric materials are expected to play a major role in the construction and assembly of a new generation of medical devices and in tissue engineering. Medical or implantable grade polymers are available in a variety of forms for producing the main body of a device (e.g. balloon catheters) or as adhesives, encapsulants, coatings, etc., to allow welding and bonding and for protecting the components of a medical device.

26.4 Surface modification techniques 26.4.1 Plasma Modification of biomaterials, including polymers, is possible using plasmatreatment processes.9 The plasma related processes are also known as gasdischarge, glow discharge polymerisation, plasma polymerisation and radio frequency glow discharge processes. A plasma may be described as a partially ionised gas composed of free radicals, ions, photons of various energy and electrons, gas atoms and molecules. In general there are two classes of plasma: cold or low temperature plasma and hot or elevated temperature plasma. A cold plasma is normally generated by low-pressure glow-discharge, whereas a hot plasma can be produced by atmospheric pressure arcs. To create a plasma, the gases (e.g. argon, oxygen, etc.) need to be excited by an outside energy (e.g. electric discharge, heat). When the surface of a polymer is exposed to a plasma, three basic reactions may occur:9 1. 2. 3.

plasma initiated polymerisation plasma state (atomic) polymerisation plasma treatment or plasma etching.

All three reactions may occur simultaneously. By altering the type of monomer or substrate used or changing the conditions of the discharge, one can manipulate the type of reaction, causing one to predominate.

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In the first process (plasma-initiated polymerisation), the plasma acts as an initiator to produce free radicals from a monomer gas (e.g. monomer containing double or triple bonds) on the surface of the substrate. This can lead to deposition of a polymer on the surface of a substrate placed inside the chamber. These reactive fragments may also undergo the second process (plasma-state polymerisation) by a recombination or other processes to form a polymer. Plasma treatment or etching is the third type of reaction occurring during a gas-discharge process. The etching plasma gases (e.g. oxygen, argon, nitrogen and water) do not polymerise but could result in a range of surface modifications of polymers including changes in surface morphology and surface chemistry. These include a different surface roughness, generation of surface functional groups, surface crosslinking, etc. The most common types of monomers used in the surface modification of biomaterials are silicone and fluorine-containing compounds.9 Silicone based polymers have a long history as medical materials and biomaterials. Fluorinated polymers also offer unique properties, such as low surface energy and inertness. Some fluorinated polymers (e.g. polytetrafluoroethylene (PTFE)) are being used in blood-contacting applications (e.g. vascular grafts) because of these properties. Examples of silicone based compounds include: hexamethyldisiloxane (i.e. [(CH3)3Si]2O) used on polyethyleneterphathalate (PET) to improve blood compatibility10 and hexamethyldisilazane (i.e. [(CH3)3Si]2NH) used for the synthesis of neurological electrodes.11 Examples of fluorine containing compounds in biomedical applications include; tetrafluoroethylene (CF2=CF2) to enhance blood compatibility of vascular grafts12 and perfluoropropane (C3F8) used in protein adsorption studies.13 A general list of plasma gas-discharge processes in biomedical applications as described by Gombotz and Hoffman is given in Table 26.1. The effects of plasma modification on the biological properties of PET were reported by Piglowski and co workers.14 PET film was modified with argon or a Table 26.1 Biomedical applications of plasma gas-discharge processes (adapted from ref. 9) Plasma treatment

Clean Sterilise Improve wettability Surface crosslinking Generate reactive sites

Plasma polymerisation and deposition Modify protein/ cell interactions

Deposit barrier films Provide reactive sites

Enhance biocompatibility Non-fouling surfaces Selective protein/ cell adsorption Cell culture surfaces

Protective coating Biomolecules Reduce leachability immobilisation Reduce sorption for therapeutics/ Control drug diagnostics delivery rate Graft polymer coating

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mixture of argon with perfluorohexane to produce hydrophilic or hydrophobic surfaces respectively. Various biological experiments in vitro and in vivo were conducted in order to study the effects of these treatments on biocompatibility. Argon plasma etching of PET films resulted in hydrophilic surfaces and the contact angle of water dropped from 70ë (before treatment) to 34ë (after treatment). The total surface energy increased from 47.4 to 65.1 mN/m. This was caused exclusively by the polar components since the contact angle of methylene oxide reagent was almost unchanged, 28ë for untreated compared to 26ë for treated. However, it was found that contact angle with water increases with storage reaching 44ë after two weeks. This was believed to be due to micromovement of functional groups and rearrangement of the upper modified surface layers with the bulk of the material. In contrast to treatment with pure argon, exposure of PET to perfluorohexane plasma resulted in a more hydrophobic surface; the contact angle of water increased to a value of 115ë. The surface energy was also low, 21.7 mN/m. In this case the surface was stable and the contact angle did not change with time.

26.4.2 Lasers The use of lasers in medical technologies and devices continues to increase whether in the area of equipment and tool manufacture, or in the area of direct interaction of the laser beam with human tissue. Currently, lasers are used for drilling, cutting, joining, hardening, surface modification and micro machining in many medical applications. The main benefits of the use of lasers in the medical sector are cleanliness, precision, lack of distortion, process speed, ease of automation, flexibility and controllability. New processes such as laser based direct metal deposition (DMD) are also starting to find applications in the field of implants. The development of new laser sources, advanced laser optics and control systems offers new opportunities for laser applications in the medical industry. The compact and energy efficient diode and Yb-YAG fibre lasers enable higher process efficiencies and are relatively easy to integrate into manufacturing lines. The development of new types of beam forming optics and beam monitoring systems, enable the process to be controlled and conducted with high precision and accuracy which are key to many medical applications. Laser types and their interaction with polymers The main commercially available laser types of interest in polymer processing are listed in Table 26.2. The different applications possible with each laser type are dependent mostly on the wavelength of light produced, which dictates the form of energy absorption in a polymer. However, for a given laser type, the processing behaviour also varies with the material type used.

0.8±1.0

0.15±0.35

Excimer

1.06

Nd:YAG

Diode

10.6

Wavelength (m)

CO2

Laser or lamp type

1,000

6,000

6,000

45,000

Max. power (W)

Reflection off mirrors

Fibre optic and mirrors

Fibre optic and mirrors

Reflection off mirrors

Beam delivery

Rectangular

Rectangular

Circular

Circular

Raw beam profile

Table 26.2 Comparison of commercially available laser and infra-red lamp sources

Complete absorption at surface in <0.5 mm Transmission and bulk heating for 0.1±10 mm Transmission and bulk heating for 0.1±10 mm Chemical bond breaking action at surface <0.01 mm

Interaction with polymers

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CO2 The CO2 laser is most commonly used for metal or polymer cutting and welding. CO2 laser radiation (10.6 m wavelength) is rapidly absorbed in the surface layers of polymers. Very rapid welding of thin polymer film is therefore possible, even with fairly modest laser powers (<1000 W). The CO2 laser beam cannot be transmitted down a silica fibre optic, but can be manipulated around a complex process path using mirrors and either gantry or robotic movement. Diode and Nd: YAG High power diode and Nd: YAG laser sources may be considered together because of the similar wavelength of light that is generated. This is in the near-infra-red range, 808±980 nm for diode sources and 1064 nm for Nd: YAG sources. At this wavelength absorption in a typical natural polymer is relatively low, but can be increased by means of additives such as pigments or fillers such as carbon black or infra-red absorbing dyes.15 The diode or Nd: YAG laser may therefore be used for welding plastic when a suitable absorbing medium is positioned at the joint line, and the beam is transmitted through the upper layer of plastic.16 The diode or Nd: YAG laser beam can be transmitted down a fibre optic enabling easy flexible operation with gantry or robot or, in the case of a diode, the whole diode laser source is small and light enough to be placed on the robot arm or gantry.17 Excimer Excimer lasers are available with average powers up to about 1 kW, but with pulse powers greater than 108 W, and focusable to very high power densities. There is a family of wavelengths available by exciting different gases within the laser. These are all in the ultraviolet, ranging from 157 nm to 353 nm, and lie in a photon energy range (3.5±7.9 eV) capable of breaking chemical bonds and splitting molecules. The C-H bond, for example, has a bond energy of 3.5 eV. Essentially, this is a cold processing technique with no material melting. This effect leads to many precision machining and surface treatment applications. Excimer laser light is absorbed by molecules on the surface of polymers (<10 m depth), and rapidly breaks the molecular bonds within the polymer structure. This leads to a rapid increase in pressure and expulsion of material over a very precise region defined by the laser beam size. The excimer laser beam is transmitted by mirrors and is often focused through masks to give the required features on the material surface.18 Excimer laser surface treatment to enhance biocompatibility The first physiological process that occurs within the initial stages of exposure is the adsorption of biomolecules onto the surface, and this is usually followed by

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cellular interactions. Both the surface topography and surface chemistry can significantly affect the type and intensity of these interactions. Illuminating many substances (in solid, liquid or gaseous form) with excimer laser photons can cause a change to either their chemical nature or structural form. Unlike many applications of lasers, ablative excimer techniques make use of several of the unique properties of this light form. The process relies on short, intense bursts of light to create a rapid rise in pressure to break bonds at the material surface. In a confined volume, the bond breaking increases the local particle number density (i.e. pressure). The corresponding rapid rise in pressure is released as a shock wave that ejects material fragments as gases and particulates at high speed. The process takes place with little excess heat transferred to the surrounding material, and as a result can be used to great effect in materials such as polymers, paper, ceramics, glasses, crystals, composite and biological tissues. Processing of materials with excimer lasers is usually most efficient when carried out using an appropriate mask inserted into the beam. Images (usually reduced) of the mask patterns are then relayed via a lens onto the workpiece. The resolution and size of the image required determines the complexity of the imaging optical system. Multi-component lenses can produce images with a resolution <1 m for highly accurate machining of small parts. As well as surface topography, the surface chemistry of polymers can be altered by photochemically induced surface reactions using excimer lasers as UV sources. Controlled surface photo-oxidation can lead to the generation of polar surface functional groups leading to changes in surface hydrophilicity and hydrophilicity. This in turn can affect the bio- and haemocompatibility of treated material. Possibilities using lasers to modify polymer surfaces to enhance biocompatibility have been investigated for the first time, jointly by Tavakoli and co-workers.19 Excimer lasers are also used in the surface treatment of polymer surfaces to improve subsequent adhesion for adhesive bonding.20,21 Certain parallels can be drawn between these two application areas. The effect of changing surface topography on neutrophil chemokinesis and fibroblast adhesion was examined using two polymeric substrates, polycarbonate and polyetherimide modified by laser treatment, to produce pillars of varying dimensions on the surfaces of these materials. The range of dimensions for the pillars were 7, 25 and 50 m square and 0.5, 1.5 or 2.5 m deep. Human neutrophils were isolated by centrifugation on Ficoll from heparinised whole blood obtained from healthy volunteers. Isolated neutrophils were exposed to the surfaces for 20 minutes and tracked using image processing and analysis techniques. The mean speed for each cell on each surface was calculated and this data statistically analysed using multivariate analysis of variance to determine any significant effect on the speed of movement due to surface topography. Confocal and SEM micrographs from typical laser treated surfaces with

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26.4 Confocal image of fibroblasts on 50 m pillars with five pulses (2.5 m deep), with material topography visible from reflection microscopy.

fibroblast cultures are shown in Figs 26.4 and 26.5. A 50 m grid (Fig. 26.4) with five pulses allowed the cells to spread. The 7 m grid with three pulses (Fig. 26.5) caused the cells to spread and elongate. The results obtained from fibroblasts demonstrated that the textured polymer surfaces showed good cytocompatibility. It was concluded that further texturing and edge effects might lead to an increase in stimulation of the neutrophils by surface treatments. In recent work, the effect of laser surface modification and the surface condition of PEEK was investigated.22 ArF laser treatment was used to treat an implantable grade of PEEK. It was shown that surface topography could influence the nature of osteoblast cell adhesion of treated PEEK. Further results of this work published more recently demonstrated that specific surface features could be produced successfully using ArF lasers.23 This treatment was compared with heat embossing. All the osteoblast-like cell types investigated attached to the PEEK surface and exhibited normal morphologies. The orientation and cell attachment of MG63 osteoblast-like cells were also affected by the type of surface feature available on PEEK.

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Figure 26.5 SEM of fibroblasts on 7 m pillars with three pulses (1.5 m deep).

PEEK is being considered as a replacement for ultra-high molecular weight polyethylene (UHMWPE) in the acetabular component of a hip replacement. PEEK has good wear resistance, which it is hoped will lead to a decrease in wear particles at the articulating surface and, therefore, a decrease in aseptic loosening. The polymer is also characterised by high strength and stiffness. In addition, sterilisation of the material is easy due to its superior resistance to chemicals, steam and moisture and ionising radiation. The biocompatibility of PEEK is dependent on its surface since it is this part of the material that comes into contact with the physiological environment. Therefore, it is hoped that modifying the surface chemistry and topography of PEEK using a technique based on the excimer laser will lead to an improved interface at the nonarticulating surface and enhance the biocompatibility of the polymer.

26.4.3 Lithography Modern microfabrication techniques, such as lithography, offer one of the most powerful tools in the precise surface engineering of biomaterials down to nano scale level.24,25 Both electron beam lithography and photolithography have been used to generate patterns and features that can be subsequently etched by either reactive ions or chemically-assisted ion beam etching. Photolithography has been used primarily in the microelectronics industry and is a method of transferring geometric shapes using a mask placed onto the surface of silicon

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wafers. The mask consists of the exact pattern, which will be transferred to the surface of the treated substance. Patterned chemical properties of dimensions in the range 10 nm±10 m could also be produced by deposition and lift-off of oxide or metal films. Microlithographic methods (e.g. photolithography) can be used to produce surface patterns. The specific surface pattern is normally made by etching into the substrate, or by deposition of a material or by chemical modification (e.g. ion implantation) through a stencil or mask as described by Gold and co-workers.24 Surface modification on the micrometre size scale can be made layer by layer, and for each layer a new pattern can be generated. Microfabrication can also be used to produce surface features of molecular size by using nanolithographic methods such as electron beam lithography and as reported by Glass et al.26 illustrated in Fig. 26.6. This will enable surface features at the nanometre scale. Brunette, Cherhroudi and co-workers27,28 were the first investigators reporting the effect of microfabricated surface topographies on biomaterials. These reports were mainly based on the effects of grooves on titanium-coated surfaces on the growth, motion and direction of cells on percutaneous and dental implants. The most commonly used microlithographic technique to produce changes in surface chemistry of materials is combining photolithography and self assembled monolayers (SAMs) typically silanes or thiols.29,30 For instance, patterned areas of one class of silicone can be generated within the matrix of a second type of silane. This will enable preferential cell attachment and spatial control of all movement and migration at surfaces. As described by Gold24 there are many possibilities for the surface modification of medical materials and biomaterials by using existing nanofabrication techniques to control molecular interactions at surfaces. A number of investigators24,31±33 used common nanofabrication techniques to control bimolecular interactions and reactions at surfaces. For instance, electron beam lithography, scanning probe microscopies as well as nanosphere28 and molecular lithographies29,30 were employed to produce surface features from submicron to nanometres scales in both inorganic and organic materials. These techniques would allow making of nanometre-sized pits on the surface of a biomaterial, which would be able to trap biochemical active enzymes or selective proteins in specified locations and desirable orientations. In this way it would be possible, for instance, to prescribe the nature of the initial protein film on an implant by positioning it in a configuration prior to insertion into the body.

26.4.4 Surface grafting Surface modification of polymers by graft polymerisation or coupling is one of the most attractive methods for producing specific surface properties including

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26.6 A scheme summarising various approaches to produce substrates with nanostructures by block copolymer micelle lithography (adapted from ref. 26). (a) Formation of extended monomicellar films and subsequent plasma treatment. Guided self-assembly of block copolymer micelles along prestructures generated by (b) photo lithography and (c) electron beam lithography. (d) Application of monomicellar films as negative electron beam resist. In (b)±(d), the complementary length scales of a diblock copolymer micelle in which a single nanodot is precisely positioned in the centre and the resolution of photo- or e-beam lithography which reaches the diameter of diblock copolymer micelle makes this concept a suitable technique for formation of different nanostructures.

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biocompatibility. The primary step in grafting is the generation of reactive functional groups on the surface of the polymer. This can be achieved by a range of chemical and irradiation techniques (e.g. UV light ionising irradiation, etc.).34 As shown in Fig. 26.7, once a reactive site is produced then selective monomers, short or long chain, low molecular weight polymers can be attached to the surface to induce specific surface properties.35±37 For instance, patterned areas of one class of silane can be generated within the matrix of a second type of silane. This will enable preferential cell attachment and spatial control of all movement and migration at surfaces. Graft polymerisation is usually based on free radical reactions as shown in Fig. 26.7. In this case the monomer reacts with the active free radical sites available on the surface of the polymer. This will lead to attachment of selective grafts to the surface of the original polymer. In graft coupling reactions, a polymer with functional end groups can be grafted using the reactive sites on the surfaces of another polymer. Hydrogel based polymers have physical properties similar to human tissue and provide a high degree of biocompatibility to the polymer surface to which they could be attached. This principle has been used by many investigators. Hydrogel forming polymers, such as poly [2-methacryloyloxy]ethyl phosphochloline (PMPC),38,39 poly[2-glucosyloxylethyl methacrylate] (PGEMA)40 poly [N-(2-hydroxypropyl)methacrylamide] (PHPMA)41 were attached to the surfaces of polyethyleneterphatalate and polyethylene using an Ar plasma induced graft copolymerisation technique. It was shown that the biocompatibility of unmodified polymers improves as a result of grafting hydrogel monomers and polymers to the surfaces of the selected substrates. In more recent work, Sugiyama and co-workers42 evaluated the biocompatibility of the surface of polyethylene (PE) films modified with various water soluble polymers, including PMPC and PGEMA, using an Ar plasma and graft copolymerisation technique. The additional hydrophilic polymers studied in this work included polyoxethylene (POE: (CH 2 -CH 2 -O) n and poly(Nisopropylacrylamide) or PNIPAAm; (CH2-CH-)n | CONHCH(CH3)2

26.7 An illustration of graft polymerisation and coupling reactions (ref. 36).

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A radio frequency power generator was used in the plasma chamber using Ar as the gas. The generator was operated at a fixed frequency of 13.56 MHz with a power output of 150 W. A schematic illustration of the plasma reactor used by Sugiyama and co-workers42 is shown in Fig. 26.8. A strip of low density PE film (7 cm  0.6 cm, Ca 30 m thickness) was washed with methanol, dried in a vacuum oven at 50 ëC, weighed and then fixed in the reactor (D in Fig. 26.8). The PE film was first exposed to Ar plasma (at 1.3 Pa with 5 W for two minutes) and then to air for one minute. The exposure of PE to air can generate surface oxidation leading to the formation of peroxides. Graft copolymerisation can be initiated by these peroxides. The PE film was then placed in a glass ampoule containing 15 mmol of monomer in 10 cm3 of absolute ethanol and allowed to undergo post polymerisation at 60 ëC for 12 hours. The grafted film was washed and soaked in methanol for 24 hours to remove any unreacted monomer or homopolymer adhered to the surface prior to testing. The biocompatibility of modified PE films was evaluated from the degree of inhibition of human thrombin activity by the addition of a synthetic chromogenic substrate (S-2238, Dai-ichi Chemicals, Japan) in the presence of 1% by weight of the relevant water-soluble polymers. This study showed that the copolymer obtained using 2(methacrylayloxy) ethylphosphorylcholine (MPC) segment exhibited an excellent biocompatibility based on the formation of a bio-membrane-like surface and because the surface adsorbed less serum protein than untreated or other PE systems.

26.8 A schematic illustration of plasma reactor (adapted from ref. 42).

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The order of biocompatibility for the copolymers studied was PMPC > PEO > PGEMA > PINIPAAM > PHPMA. It was found that PMPC-grafted PE suppressed the amounts of proteins such as albumin, -globulin and fibrinogen more than other PE systems because the PMPE segments had the largest amount of free water around the main chain. In an earlier study, Sugiyama and co-workers38 investigated grafting a number of vinyl monomers, including MPC on a PET film to enhance biocompatibility. Again graft copolymerisation was carried out using a low temperature, Ar plasma, post polymerisation technique. The contact angle of the modified PET film decreased from  ˆ 68ë (original PET film) to  ˆ 26ë for PMPC-grafted PET. The modified PET films absorbed less serum protein than the original PET film. Biocompatibility was also evaluated by the cellular response to the modified PET film using mouse fibroblast cells. It was found that cell adhesion did not take place on PMPC grafted PET. Mirzadeh and co-workers used CO2 laser as a source of infra-red energy to initiate grafting of acrylamide (AAm) and 2-hydroxyethyl methacrylate (HEMA) to the surface of a thermoplastic elastomer, ethylene propylene rubber (EPR).43,44 Infra-red spectroscopy, electron microscopy and energy dispersive X-ray analysis (EDXA) were used to characterise the modified polymer. The grafted poly (AAm) and poly (HEMA) were found to have a fractal morphology. Fractal surfaces provided both hydrophilic and hydrophobic sites, making EPR especially suitable as a biomaterial.

26.4.5 Coatings There are many different coatings that are available which could be deposited on the surfaces of medical materials and biomaterials to produce specific surface properties Acrylic based polymeric bone cements are being used in hip replacements and other orthopaedic surgical applications to assist bone ingrowth and implant fixation. Echogenic polymer coatings have also been developed to enhance ultrasound visibility of medical devices, particularly biopsy needles.45 The biomedical coatings have been the subject of many recent reports.46,47 Coatings used in short-term applications include those used in single-use and disposable devices (e.g. coated catheters, sensors, and biopsy needles). Coatings used in long-term applications include those applied in implants (e.g. heart valves, pacemakers and prosthetics). Biocompatible polymer coatings are also being used in drug-eluting stents.48 Coronary stents are used in interventional cardiology for the treatment of complex lesions and multivessel coronary disease. However, the related procedure may become complicated due to a clinical problem called in-stent restenosis. This is caused by neointinal proliferation, which starts a few days after stent implantation as a general vascular response to injury. One approach to treat restenosis is local stent delivered drug therapy. The polymeric coating on

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26.9 An example of antimicrobial coating on a surface of a catheter (adapted from ref. 49). MPEG = methacrylate polyethylene glycol; 2000 and 350 are molecular weights; MA = methacrylate.

the stent can carry an antirestenotic drug, which could be released in a controlled manner at the required and specified location. The polymer coatings for these applications have to be biocompatible, haemocompatible and have chemical reactivity with the selected drugs. Typical examples of polymeric systems used as drug delivery coatings on stents include polyurethanes and polysiloxanes. Polymer-based coatings are also applied to enhance thrombo-, antimicrobial resistance and increase lubricity as well as biocompatibility (Fig. 26.9). Another example of coating available for medical application is Parylene. Parylene is a pure, crystal-clear, polycrystalline, and amorphous linear polymer that represents a viable option for medical coating applications because of its biocompatibility and biostability. A transparent Parylene film is formed from a pure molecular precursor (a monomer gas), ensuring that the finished film has no contaminating inclusions. Because the conversion from monomeric gas to polymer film is direct, no solvents, plasticisers, catalysts, or accelerators are used. The resulting film has low potential to trigger an immune response. Parylene has been shown to be resistant to the effects of body fluids, electrolytes, proteins, enzymes, and lipids. The film also forms an effective barrier against the passage of contaminants from a coated substrate to the body or surrounding environment. Some suppliers of Parylene coating will have both drug and device master files with the FDA for the Parylene polymers. This information is available on request, and includes data on body-tissue and bloodcompatibility studies

26.4.6 Haemocompatibility During the last thirty years considerable research activity has been devoted to studying blood and biomaterial interactions in an effort to design and produce blood compatible (haemocompatible) materials. The application of biomaterials in blood contacting devices has significantly increased in recent years. Typical

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examples of devices where polymers are used as a part or whole of the blood contacting device include; catheters (e.g. PET based balloon catheters used in cardio-angioplasty), blood delivery tubes (e.g. PVC tubing) or artificial vascular grafts used in vascular surgery (e.g. fluoropolymers as large vessel replacements). Recently some polymers (e.g. polyurethanes) are also being considered as part of an artificial heart, particularly heart valves. As described by Sandhu and Luthra,49 there have been two main approaches to provide or enhance haemocompatibility: · development of nonthrombogenic materials · development of antithrombogenic materials. A number of hydrophilic polymers have shown nonthrombogenic properties such as polyethylene glycol/polyethylene oxide and polyvinylpyrrolidene and phosphorylcholine (PCe) based copolymers. Nonthrombogenic polymers are those that do not activate cells or platelets or enhance protein binding. It is believed that polymers containing PCe surface functional groups will provide nonthrombogenic characteristics. This is based on the fact that PCe head groups are available on the surface of red and white cells and it was proposed that PCe mimics the outer cell wall (see Fig. 26.10). The second main method of generating a haemocompatible surface is to use heparin or heparin derivatives on antithrombogenic materials. Heparin performs by binding to antithrombin.49 This complex catalyses the binding and inhibition of thrombin thereby preventing clot formation; the bonding of heparin has increased the inhibitory effect on antithrombin or thrombin. Heparin and heparin-like substances will function by providing an antithrombogenic charac-

26.10 Phosphorylcholine head group and polymers.

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26.11 The nonthrombogenic/antithrombogenic polymer's triple endotheliallike action for biocompatibility (adapted from ref. 49).

teristic. Heparin acts by binding to antithrombin and therefore preventing clot formation. Sandhu and Luthra49 have also reported the development of a new haemocompatible material. This material is said to provide both antithrombogenic and nonthrombogenic properties which provide an endothelial-like action that prevents protein adsorption and inhibits thrombin at the same time. The new nonthrombogenic/antithrombogenic polymer coating can provide triple endothelial-like action.50 It does this through three components: heparin, a negative charge and hydrophilicity as shown in Fig. 26.11. Nonleaching heparin molecules are covalently bonded onto the surface to provide similar beneficial effects to heparan sulphate in the natural endothelium. Sulphate and sulphonate groups, which carry a strong negative charge, are incorporated into the functional layer of the material to repel blood cells and proteins. Heparan sulphate in the vascular endothelium is similarly a heavily sulphated molecule with a negative electrical charge, as is heparin in the novel material. Mimicking biologically inert membrane surfaces has been one of the most active areas of recent research. As described earlier, phosphorylcholine (PC) groups are the most prominent polar head groups present in natural membranes. The interactions between blood and the surfaces of biomaterials have been described in a number of publications.51±55 When blood contacts a foreign body which may have a thrombogenic surface, the haemostatic system, particularly the coagulation process, is instantly initiated. Processes occurring as a result of these interactions were described by Yianni,51 as shown in Table 26.3.

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Table 26.3 Processes occurring as a result of contact between blood and a thrombogenic material (adapted from ref. 51) Protein absorption

Platelet interaction

Activation phenomena

Fibrinogen Immunoglobulin Alumin C-reactive protein

Adhesion Activation

Clotting cascade in particular Factor XII and complement cascade (C3 to C3a)

The thrombogenic surface rapidly acquires a layer of adsorbed plasma proteins, the nature of which mediates the following thrombotic events. At least 60% of this adsorbed protein layer is fibrinogen, which is present in blood (2.4 mg/ml of blood). The structural characteristics and limited solubility of fibrinogen were thought to predispose it to binding to the surface of foreign materials. Following adsorption of fibrinogen, platelet adhesion and activation occurs. The combination of platelet adhesion and activation and fibrinogen binding and activation of the clotting factor cascade leads to the formation of clots at the surface of a material with thrombogenic characteristics. Forbes52 described various processes occurring within minutes, days and months of exposing a material to blood. These are listed in Table 26.4. The main events occurring within minutes include protein adsorption, cell adhesion and inflammation. These lead to events such as thrombus formation and fibrinolysis. The longer exposure of a biomaterial to blood may lead to embolisation, calcification and changes in material properties. Covalent bonding of phosphorylcholine functional groups to a number of commercially available polymers to enhance their blood compatibility was also reported by Chapman and Durrani.56±58 Reactive compounds were prepared which allowed attachment of PCe units to surfaces of polymers containing carboxyl or hydroxyl groups. The modified polymer retained its mechanical properties, but exhibited the surface characteristics of all membrane surfaces. As mentioned earlier, this method to improve blood compatibility is based on low levels of interaction between PCe and the surfaces of biomaterials and proteins Table 26.4 The effect of time and sequence of events with artificial surface (ref. 52) Minutes

Days

Months

Water and ion interactions Protein adsorption Cell adhesion Local fibrin deposition Inflammation Embolisation

Changes in proteins Continued cell adhesion Thrombus formation Fibrinolysis Chronic inflammation Embolisation

Embolisation Calcification Changes in material properties, e.g., fatigue

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and cells in blood. More recently, attachment and coating of PCe polymers to polyvinylchloride (PVC) and polyethylene (PE) were reported by Yianni59 and Liu and co-workers.60 PVC is one of the most commonly used polymers for disposable medical devices and there are many applications where direct contact with blood will be required. It was described59 that PVC could be typically coated with methacryloylphosphorylcholic (MPC)/laurymethacrylate (LM) copolymer. The PCe polymer was dissolved in a suitable solvent such as ethanol and then passed through a 5 m filter, strips of PVC were then dipped into this solution. The coated PVC was then air dried before being evaluated by haemocompatibility testing. The results obtained clearly demonstrated that platelet adhesion and activation were greatly reduced when PVC was coated with a PCe polymer. It was concluded that the haemocompatibility of PVC could be significantly enhanced by coating its surface with phosphorylcholine-containing polymers. The PVC coated with PCe polymer was found to be stable and could resist degradation during sterilisation. All toxicological regulatory requirements (ISO 10993 Part 1) were also met by PCe polymers and PCe-coated devices. In an earlier report Liu and co-workers60 described the surface modification of PCe membranes using PCe derivatives to enhance platelet compatibility. Acrylic acid (AA) was graft-copolymerised onto the surface of PE by UV irradiation. Prior to this process, PCe film was immersed in an aqueous solution containing a photo-initiator, sensitiser and various organic solvents. PCe with various spacer lengths was introduced onto the surface of PE by a series of chemical reactions. Ethylene glycol (EG), butanediol (BDO), polypropylene glycol (PPG) and polytetramethyl glycol (PTMG) were used as spacers. The platelet compatibility of PCe-modified film was evaluated by a platelet adherend test. This work showed that the platelet compatibility of the PE film was affected by the existence of various functional groups on the film surface. It was found that the amount of adhered platelets decreased in the order: PCe-POH (phosphoryl oxychloride (water treated PE film)) > PE-PTMG > PE-AA > PE-PCe. The length of lipophilic spacer between the PCe groups and the PE surface will affect the stability of the film surface.

26.5 Future trends Interest in developing novel surface modification techniques to induce specific surface properties is expected to be an active area for new research and development. As well as improving biocompatibility, the ability to induce additional and complementary surface properties such as haemocompatibility, permeability, microbial resistance, wear resistance and lubricity of polymers will continue to be of major interest to the biomaterial community and device manufacturers. Achieving a high degree of biocompatibility and unique surface properties will lead to a new generation of materials for applications in both

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short- and long-term implantable devices, which can provide satisfactory performance in specific applications in contact with cells, tissue or blood. However, one of the challenges to using nanotechnology and nanofabrication techniques for the surface modification of biomaterials has been the ability to produce large quantities of modified surfaces with consistency. Direct surface modification techniques, which can consistently and effectively modify polymer surfaces at a low or reasonable cost and can be adapted both for small and large surfaces as well as to a complicated geometry will become very important. These are expected to be favoured by industrial organisations involved in supplying medical and biomaterials or in the manufacturing of short- and longterm implantable devices.

26.6 Acknowledgements The author wishes to thank Professor Pankaj Vadgama for giving him the opportunity to prepare this chapter. Thanks are also due to Mrs Sue Dunkerton, Director of Medical Devices, Faraday Partnership for her helpful comments and Mrs Janette Whiting for typing this chapter. Finally he would also like to thank his wife for all her support during the preparation of this chapter.

26.7 References 1. Hoffman A S: Present and emerging applications of polymeric biomaterials, Clinical Materials 1992, 11, 13±18. 2. Griffith L G: Polymeric biomaterials, Acta Materialia, 1 January 2000, Volume 48, Issue 1, 263±277. 3. Chiellini E, Sunamoto J, Migliaresi C, Ottenbrite R M and Cohn D: Biomedical polymers and polymer therapeutics, Kluwer Academic Publishers, Dordrecht, 2001, pp. xviii, 451. 4. Dittgen M, Durrani M and Lehmann K: Acrylic polymers, a review of pharmaceutical applications, Journal of STP Pharma Sciences 1997, 7 (6), 403±437. 5. ISO 10993, Biological Evaluation of Medical Devices www.iso.ch/iso/en/standards. 6. Williams D F: Williams dictionary of biomaterials, Liverpool University Press UK, 1999 ISBN: 0853237344. 7. Williams D F: Revisiting the definition of biocompatibility, Medical Device Technology Magazine, December 2003. 8. Williams D F: Revisiting the definition of biocompatibility, Medical Device Technology Magazine, March 2004. 9. Gombotz W R and Hofman A S: Gas discharge techniques for biomaterial modification, CRC Critical Reviews in Biocompatibility, 1987, Volume 4, Issue 1, 1±42. 10. Yasuda H and Bumgarner M O: Improvement of blood compatibility of membranes by discharge polymerisation, Permeability of Plastic Films and Coatings to Gases, Vapours and Liquids. Hopfenberg H B ed. Plenium Press, New York, 1974, 453. 11. Cannon J G, Dillon R O, Bunshah R F, Crandall P H and Dymond A M: Synthesis of

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Surfaces and interfaces for biomaterials a fine neurological electrode by plasma polymerisation processing. Journal Biomedical Material Reviews, 1980, 14, 279. Garfinkle A M, Hofman A S, Ratner B D, Reynolds, L O and Hansen S R: Effects of tetrafluoroethylene glow discharge on patency of small diameter vascular grafts. Trans. Am. Soc. Am. Artif. Intern. Organs. 1984, 30, 432. Hague H and Ratner B D: The preparation of plasma deposited films with surface energies varying over a wide range. Journal of Applied Polymer Science, 1986, 32, 4369. Piglowski J, Goncarz I, Stainszewska-Kus Plauch D, Szmonowicz M and Konieczny A: Influence of plasma modification on biological properties of polyethyleneterephthalate. Biomaterials 1984, Vol 15, No 11, 909±916. Jones I A, Hilton P A, Sallavanti R and Griffiths J: Use of infrared dyes for transmission laser welding of plastics, Proc. ICALEO, Nov 1999. Potente H, Korte J and Stutz R: Laser-transmission welding of PE-HD, Kunstoffe, 1997, 87, 3, 348±350. Hug M and Rudolf T: Assessment of different high power diode lasers for material processing, Proceedings: Lasers in Materials Processing (Munich) June 1997. Gower M C: Excimer lasers: principles of operation and equipment, and current and future application in industry and medicine. Laser Processing in Manufacturing edited by Crafer R C and Oakley P J, pp 163±261, Chapman and Hall, 1993. Hunt J A, Tavakoli S M, William R L and Riches S T: Laser surface modification of polymers to improve biocompatibility. 12th European Conference on Biomaterials, Porto, Portugal, 10±13 September 1995. Tavakoli S M and Riches S T: Laser surface modification of polymers to enhance adhesion, Part 1 ± polyolefins, Proceedings of ANTEC `96, 5±10 May 1996, Indianapolis, USA. Tavakoli S M and Riches S T: 'Laser surface modification of polymers to enhance adhesion Part II ± PEEK, APC-2, LCP and PA, ANTEC 2000 Conference, 7±11 May 2000, Orlando, Florida, USA. Corfield V I, Tavakoli S M, Brooks R, Walton C E, Cameron R E and Bonfield W: Surface modification of PEEK to enhance biocompatibility. 17th European Conference on Biomaterials 11±14 September 2002, Spain. Corfield V, Snelling H V, Cameron R E, Tavakoli S M and Bonfield W: ArF laser ablation of PEEK to introduce microscopy and control cell interaction, 7th World Material Conference 17±21 May 2004, Sydney, Australia. Gold, J: Microfabrication for biological applications: Preparation, characterization and biological evaluation, Doctoral thesis, Dept of Applied Physics, Chalmers University of Technology, Gothenburg, Sweden, 1996. ISBN 91-7197-275-7. Clark P: Cell behaviour on micropatterned surfaces, Biosensors and Bioelectronics 1994, 9 657±661. Glass R, MoÈller M and Spatz J P: `Block copolymer micelle nanolithography'. Nanotechnology 2003, 14, 1153±1160, Institute of Physics Publishing. Brunette D M, Kenner G S and Gould T R L: Grooved titanium orient growth and migration of cells from human gingival explants, J Dent Res 1983, 62, 1045±1048. Chehroudi B, Gould T R L and Brunette D M: Titanium-coated micromachined grooves of different dimensions affect epithetial and connective tissue cells differently in-vivo. Journal of Biomaterial Materials Reviews, 1990, 24, 1203±1219. Whitesides G M: Self-assembling materials, Sci Am 1995, 273, 114±117.

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30. Nuzzo R G and Allara D L: Adsorption of biofunctional organic disulfides on gold surfaces, Journal American Chemistry Society, 1983, 105, 4481±4483. 31. Hulteen J C and van Duyne R P: Nanosphere lithography, a materials general fabrication process for periodic particle array surfaces, J Vac Sci Technology A 1995, 13, 1553±1558. 32. Douglas K, Clark N A and Rothschild K J: Nanometer molecular lithography, App Phys Lett 1986, 48, 676±678. 33. Pum D, SaÂra M, Messener P and Sleytr U B: Two-dimensional (glyco) protein crystals as patterning elements for the controlled immobilization of functional molecules, Nanotechnology 1991, 2, 196±202. 34. Garbassi F, Morra M and Occhiello E: Polymer surfaces from physics to technology, John Wiley & Sons, Apr 1988, Chapter 7, pp 272±284. 35. Uyama Y, Kato K and Ikeda Y: Surface modification of polymers by grafting. Advances in Polymer Science, 1998, 137, 1±39. 36. Ikada Y: Surface modification of polymers for medical applications, Biomaterials, 1994, 15 (10), 725±736. 37. Prime K L and Whitesides B M: Self-assembled organic monolayers; model systems for studying adsorption of proteins at surfaces, Science 1991, 252, 1164±1167. 38. Sugiyama K, Kato K, Kido M, Shiraishi K, Ohga K, Okada K and Matsuo O: Grafting of vinyl monomers on the surface of a poly(ethylene terephthalate) film using Ar plasma-post polymerization technique to increase biocompatibility, Macromolecular Chemistry and Physics, 199 (6), June 1998, pp 1201±1208. 39. Iwasaki Y, Mikami A, Kurita K, Yui N, Ishihdra K and Nakabayashi N: Reduction of surface-induced platelet activation on phospholipid polymer. Journal of Biomedical Materials Research 1997 36 (4), pp 508±515. ISSN: 0021-9304. 40. Chen Z, Zhang R F, Kodama M and Nakaya T: Novel grafted segmented polyurethane-bearing glucose groups, J Biomaterials Science, Polymer edn, 1999, 10 (9), pp 901±916, ISSN: 0920-5063. 41. Sugiyama K, Mitsuno S and Shiraishi K: Adsorption of protein on the surface of thermosensitive poly (methyl methacrylate) microspheres modified with the N-(2hydroxypropyl)methacrylamide and 2-(methacryloyloxy)ethyl phosphorylcholine moieties, Journal Polymer Science Part A, Polymer Chemistry, 1997, 35, 16, pp 3349±3357. 42. Sugiyama K, Matsumoto T and Yamazaki Y: Evaluation of biocompatibility of the surface of polyethylene films modified with various water soluble polymers using Ar plasma-post polymerisation technique, Macromolecular Materials Engineering 2000, 282, 5±12. 43. Mrizadeh H, Katbab A A and Burford R P: CO2-pulsed laser induced surface grafting of acrylamide onto ethylene-propylene rubber, International Radiation Physics Chemistry, 1993, 41, 507±19. 44. Mrizadeh H, Khorasani M T, Katbab A A, Burford R P, Soheili Z, Golestani A and Goliaei B: Biocompatibility evaluation of laser-induced AAm- and HEMA-grafted EPR.Part 1: In-vitro study, Biomaterials, 1995, 16, 641±8. 45. Tavakoli S M, Kellar E J C, Nassiri D and Joseph A E: A novel polymeric coating for enhanced ultrasound imaging of medical devices, Journal of Applied Medical Polymers, December 2002. 46. Biocompatible Coating The Impact of Surface Technology on Medical Markets, Technical Insights, John Wiley & Sons Inc, June 2000 (see the summary at

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www.mindbranch.com). 47. Medical Coatings, Technology Assessment Associates, March 2002 (see summary at www.tech-assessment.com). 48. Al-Lamee K and Cook D: Polymer coating techniques for drug-eluting stents, Medical Device Technology, January 1, 2003. 49. Sadhu S and Luthra A: New biointeracting materials, Medical Device Technology, October, 2002, 13 (8), pp 10±16. 50. <www.medtronic.com> The nonthrombogenic and antithrombogenic polymer is commercialised by Medtronic Inc. under the name of Trillium and is licensed for its cardiopulmonary bypass products. 51. Yianni Y P: Biocompatible surfaces based upon biomembrane mimicry structural and dynamic properties of lipids and membranes, Portland Press Research Monograph, eds P J Quinn and R Cherry 1992, 187±215. 52. Forbes C D and Courtney J M: Blood material interactions ± implications for clinicians, Scott Med J. August 1995, 40 (4), pp 99±101. 53. Courtney J M and Forbes C D: Thrombosis on foreign surface, British Medical Bulletin, 1994, 50 (4), 966±981. 54. Courtney J M, Sundaram S, Yin H Q and Forbes C D: Artificial surfaces and blood interactions, Vascular Medicine Review, 1994, 5, 42±49. 55. Courtney J M, Lamba N M K, Sundaram S and Forbes C D: Biomaterials for bloodcontacting applications, Biomaterials, 1994, 15 (10), 737±744. 56. Chapman D and Durrani A A: European Patent Application No 8151300 356-4 January 1985. 57. Durrani A A, Hayward J A and Chapman D: Biomembranes as models for polymer surfaces II: the synthesis of reactive species for covalent coupling of phosphorylcholine to polymer surfaces, Biomaterials, 1986, 7, 121±125. 58. Chapman D and Charles A C: A coat of many lipids in the clinic, Chemistry in Britain, March 1992, 253±256. 59. Yianni J P: Making PVC more biocompatible, Medical Device Technology, September 1995, 20±29. 60. Liu J H, Jen H L and Chung Y C: Surface modification of polyethylene membranes using phosphorylcholine derivatives and their platelet compatibility, Journal of Applied Polymer Science, 1999, 74, 2947±2954.

27

Issues concerning the use of assays of cell adhesion to biomaterials S L J A M E S and S M I K H A L O V S K Y , University of Brighton, UK, P V A D G A M A , University of London, UK and P E T O M L I N S , National Physical Laboratory, UK

27.1 Introduction Cell adhesion, either to other cells or to biological substrata such as basement membranes, is fundamental to the organisation of tissues in multicellular organisms. Furthermore, bacterial pathogenicity in many species of bacteria relies, in part, on the attachment of these pathogens to cellular surfaces. Such cells have thus evolved sophisticated and subtle mechanisms to effect (and occasionally prevent) such adhesion, involving chiefly the binding of cellular adhesion molecules to substrata ligands. Also, it must not be overlooked that a cell membrane, even one devoid of adhesion molecules, will interact with an adjacent surface via reasonably well-defined physico-chemical interactions as observed, say, in colloid science. Such interactions can, of course, be either attractive or repulsive. The introduction of a non-biological biomaterial into this environment presents such cellular systems with a situation they have not fully evolved to cope with, though the foreign body response is an established reactive cascade designed to degrade and possibly partition off a non-biological intrusive component. Nevertheless, such reactions are likely to be poorly predictable (see, for example, Hu et al., 2001). In many instances, cellular adhesion to a biomaterial is of critical importance to the initial success or failure of the material's application. These applications range from, say, indwelling vascular or urinary catheters, where adhesion of either host cells or bacteria is unwanted, through to implanted prostheses, where cellular adhesion to the prosthetic device might be either desired, as in many orthopaedic devices, or not, as in some heart valves. At the other extreme of tissue engineering scaffolds, cellular adhesion may be largely desired. Methods allowing the measurement (either quantitatively or semi-quantitatively) of the adhesion of cells to artificial substrates therefore are clearly needed, but because of the very artificial nature of the system, considerable thought must be given to the applicability and design of the adhesion assay, and the interpretation of results. Obvious areas which should be addressed include:

746 1. 2. 3. 4.

Surfaces and interfaces for biomaterials clearly defining the objectives of the assay within the context of what the investigator actually requires ensuring that the assay is yielding results which are measuring parameters the investigator intended to measure ensuring that both the biomaterial surface and the cells used as adhesion probes can be produced in a reproducible and relevant way choosing the type of assay most relevant to the system being tested.

Consideration is given to each of these areas in the sections below. It should not be overlooked that adhesion is not the only measurable which might indicate the success or failure of the interaction of a biomaterial with cells. For instance, changes in cell morphology, such as spreading, are considered important (e.g. Acaturk et al., 1999; Hallab et al., 1995), as are metabolic responses of the cells, such as expression of membrane antigens and other molecular markers (reviewed by Hunt et al., 1997). The proliferation and differentiation of cells seeded onto biomaterials can be measured (e.g. Deligianni et al., 2001), as can the death of cells in response to toxic components of the biomaterials (e.g. Cenni et al., 1999). Of considerable importance in the assessment of tissue engineering scaffolds, assays for the estimation of cell migration through the scaffold are being developed (e.g. Gosiewska et al., 2001).

27.2 Measurement objectives Because of the wide range of applications of biomaterials to biomedical problems, it is clear that different investigators will consider different objectives important when measuring cellular adhesion. Probably the major division of adhesion assays is between those which measure the adhesion of a population of cells to a surface, with the measured parameter usually being the number of cells left adhering to the surface after an attempt at their removal (e.g. Reyes and Garcia, 2003), and those which measure the adhesion of a single cell to a surface, with the measurement usually being of the force required to remove the cell from the surface (e.g. Huang et al., 2003). Either of these approaches is valid depending upon the requirements of the investigator.

27.2.1 Cell population techniques Cell population based assays tend to be cheap, because expensive equipment is often not required, and they also average out variations in cell-to-substrate adhesiveness. This variation arises both because of variations in biomaterial surface properties, and variations in cell phenotype used as the probe (see below). Relatively simple statistical analysis can then provide a usable index of that biomaterial's adhesiveness for those cells employed.

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Perhaps less obvious, some authors have suggested that cell adhesion to substrata is a stochastic process (reviewed by Cheng et al., 2000) arising out of the suggestion that each cell binds to a protein-adsorbed surface through a small number of receptors, the total number bound varying with time. Again a cell population based assay tends to have an advantage because most such assays average cell adhesion over time as well as area. However, a number of disadvantages of such assays present themselves. Adhesion versus de-adhesion Firstly, the investigator needs to be clear that it is cell de-adhesion that is to be studied, and not cell adhesion. This is because most population based assays rely upon cells being allowed to adhere to a material surface, and then a shear or normal force applied to remove them, the number retained at a particular force, or the force required to remove a predetermined fraction of the cells being the measured outcome. This is often reported as being a measure of cell adhesion, but this assumes adhesion and de-adhesion are similar but complementary processes. This may not be correct. For instance, Chen and Springer (2001) have investigated the attachment of neutrophils to a P-selectin coated polystyrene surface in an environment in which the viscosity of fluid flowing over the cells, in a parallel sided flow cell, was varied by the addition of Ficoll. This allowed independent variation of shear rate and shear stress. It is clear from this work that attachment, at least by tethering, is defined by shear rate, whereas dissociation is shear stress defined, fitting the Bell model (Bell, 1978) better than others available. This work relates to only one receptor-ligand pair, one particular flow regimen, and tethered attachment rather than rolling attachment. However, these authors helpfully interpret their findings by suggesting that attachment is a function of translational velocity of the cell relative to the surface, which would indicate that the collision between receptor and ligand can be of such short duration that a stable bond does not form (hence shear rate dependence). Detachment, however, implies that such bonds have already formed, and thus the force of detachment is related to bond strength (hence shear stress dependence). It is reasonable to suppose that this principle might also apply to other cell interactions with biomaterials. It is interesting to note that these authors (Chen and Springer, 2001) further developed an adhesion, as opposed to de-adhesion assay, by monitoring transient adhesions and tether formation by videomicroscopy in a flow cell. This may well be the ideal method for measuring cell attachment to a biomaterial using a cell population approach. Mode of detachment A further issue is that the exact mode of detachment cannot be controlled. At least two clearly differentiated modes of detachment have been widely cited in

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the literature. Since the early days of interest in cell adhesion to biomaterials, driven by interest in erythrocyte adhesion to putative prosthetic heart materials and materials used in extra-corporeal circulation, the phenomenon of cell tethering has been observed (e.g. Kochen, 1966). Here, part of the cell membrane attaches to the biomaterial surface and under shear conditions becomes pulled out into a long membranous process or tether. Cell detachment in shear flow is most likely to occur as a result of tether rupture, but the tether may also detach from the surface. This phenomenon has now been observed in cells other than erythrocytes, largely investigated by Hochmuth and Mohandas (1972). Both theoretical considerations (Cheng, 2000) and experimental observations, albeit in the slime mould D. discoidium (Decave et al., 2002; Garrivier et al., 2002) investigated the removal of cells from a material during shear flow by peeling. Here, a relatively small number of ligand-receptor bonds needed to be broken per unit time, but the shear force on the cell had to be maintained for a sufficient period for the cell to peel completely from the surface. Other modes of detachment almost certainly also occur, for instance, both tethering and peeling relate to removal by shear. A number of tests of adhesion, in particular centrifugation based tests, rely on a force that is normal to the cell being applied and this pattern of ligand rupture is quite different from shear removal. Defining shear stress close to the surface A third difficulty is the necessity to define shear conditions very close to the biomaterial surface, in a shear field, such that the actual shear stress acting on cells can be accurately determined. This requires respectively the determination of boundary layer behaviour close to the surface of the material being tested, of the influence of attached cells on the flow over cells immediately `downstream' of them, and of the shape and degree of protrusion of cells into the flowing stream. Indeed, even computation of the shear stress acting on a single cell in a flow chamber relative to inlet flow rate is not simple (see, for example, Brooks and Tozeren, 1996).When considering an adhesion, rather than a de-adhesion assay, the exact behaviour of cells flowing close to a surface should be considered. Decades ago, Goldman et al. (1967) developed a model for the flow of rigid spheres in simple laminar shear flow near a surface, which suggested a separation between such spheres and the surface. Indeed soon afterwards, it was widely reported and experimentally observed that red blood cells flowing in a narrow tube created a cell-poor layer immediately adjacent to the tube wall, the so-called plasma skimming layer (Charm and Kurland, 1974). More recently, Tempelman and coworkers (1994) have studied the behaviour of rat basophilic leukaemia cells flowing in laminar shear flow close to a surface, in comparison

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with the flow of the same cells fixed in glutaraldehyde and of rigid spheres. They found that the unfixed cells travelled faster than either the fixed cells or rigid spheres, and that application of Goldman's model suggested the cells were separated from the surface by at least 550 nm, a distance too great for receptor mediated adhesion (tens of nanometers). The very fact that adhesion is observed suggests that the cells approach the surface more closely, probably, the authors suggest, a result of cell deformability and cell surface roughness. It is thus clear that events close to the surface in a laminar shear field are complex and not easily modelled due to the constant changes in cell shape.

27.2.2 Single cell techniques Single cell techniques usually employ either a micropipette, a microcantilever or an optical device to remove a single cell from a surface. This has the advantage of generating data relevant to a single cell-surface event, and can sometimes detect individual ligand-surface interactions. It also provides truly quantitative measurements of the force required to remove a single cell from a substrate, a parameter more closely related to established concepts of adhesion. Disadvantages here are essentially the converse of those of the population approach, i.e., variations are not averaged out, and sophisticated equipment such as an atomic force microscope (AFM) is required. As with population assays, a single cell may be removed either with normal force (e.g. Thie et al., 1998) or with lateral force (e.g. Yamamoto et al., 1998). Problems of interpreting normal force removal are addressed in the next section, but removal in the lateral mode, usually with a microcantilever, introduces the problem of cell deformation and possibly even damage.

27.2.3 Sensitivity of the adhesion assay A valid question which should be raised in any discussion of measurement objectives is, what assay sensitivity is required? This is often overlooked in considerations of cell adhesion assays, the implicit assumption being that sensitivity should be as great as possible within the confines of the technique. Unfortunately, this might mean that considerable effort is expended in achieving assay sensitivity far in excess of that needed for a particular application, or alternatively, even the highest sensitivity achievable by a given method may be insufficient for a specific application. Many techniques involving both single cell and cell population approaches concentrate upon defining exactly what forces are needed to remove cells from a surface. It may be argued that in an industrial context, where the objective may be to compare the cell adhesiveness of several biomaterial surfaces, good repeatability of an assay may be more important than obtaining absolute values of binding forces. Thus a simple assay such as that used by Wilkins et al. (1990)

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for measuring adhesion of bacteria to IUD threads may be appropriate. Here the thread was simply shaken with buffer for a predetermined time and the number of bacteria remaining on the thread counted. This approach shows good reproducibility, but cannot determine actual binding forces. From the value of this simple technique, it could be argued that some assays merely need to determine whether cells attach to a surface or not. Thus assays such as those used by Acaturk et al. (1999) in which cells were allowed to settle on a surface then simply rinsed to remove cells which had not adhered, yielded a useful `allor-nothing' type of assay. An extreme of this argument might be that such assays could be completely irrelevant, because we usually have no idea what forces cells adhering to biomaterials are subjected to in vivo. Cells adhering to catheters in blood vessels may be subjected to relatively high shear stresses, whereas those adhering to, say, artificial skin scaffolds may be subjected to virtually no removal forces at all. This rather misses the point that such assays actually measure the potential for interaction of a cell with that surface, not necessarily the forces needed to remove them from the surface Alternatively, studies considering the actual mechanisms of attachment of cells to biomaterials would benefit from assays able to measure adhesion forces of the order of those between a single receptor and ligand. Here, very high sensitivity and the generation of absolute force values is important. A usable, if not particularly elegant indicator of sensitivity might be that the assay employed for a particular purpose must be sensitive enough to just distinguish between two materials which differ by an amount which is industrially, medically or academically important.

27.3 Issues of interpretation of adhesion measurements Many cell adhesion assays measure the force needed to remove a cell or cells from a surface. However, less attention than is deserved has been devoted to the mechanism of detachment. Not only is the issue of peeling as opposed to tethering detachment important, as mentioned above, but the location of adhesion failure affects the interpretation of assay results directly. Failure of adhesive bonding has long been recognised to occur not only through the adhesive separating from the bonded material, but possibly by mechanical failure of the adhesive or material itself. This type of analysis has rarely been applied to cell de-adhesion assays.

27.3.1 Cell bonding to biomaterial surfaces The initial stages of a cell bonding directly to a perfectly clean biomaterial surface is achieved largely via long-range non-biological physicochemical

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interactions such as van der Waals and electrostatic interactions, in turn influenced by material surface energy and surface charge density (Razatos et al., 1998a). This has been studied more extensively in bacteria than in eukaryotic cells, although mechanisms of attachment may differ. In bacterial cells, this initial reversible stage may, if the cell can overcome energy barriers and approach the surface closely, translate into an essentially irreversible interaction mediated via the production of an extracellular matrix and a number of bacterial adhesins. (Costerton et al., 1995). For example, Razatos and co-workers (1998a,b) have investigated the forces which exist between an AFM cantilever tip and a lawn of E. coli, and by modifying the tip with a biomaterial of interest, were able to determine whether the initial forces existing between the bacteria and biomaterial were attractive or repulsive. Furthermore, by using strains of E.coli differing subtly in surface structure, they were able to show that the adhesion force was primarily influenced by the bacteria's core lipopolysaccharide and the production of the capsular polysaccharide. It may be reasonably argued that if one is considering human cells interacting with a putative implant material, this scenario is irrelevant anyway. Early work by Vroman and colleagues (1971) showed that when biomaterial surfaces are exposed to biological fluids containing proteins, these proteins are adsorbed onto the material surface in a particular way. More recently, it has become clear that this affects cell adhesion to those surfaces because such proteins contain peptide motifs which are ligands for adhesion molecules found on cell surfaces. The most widely studied of such motifs is the RGD (arginine, glycine, aspartic acid) tripeptide which binds to some 11 of the 24 known alpha-beta integrins on cell surfaces (for a recent review of RGD modified polymers see Hersel et al., 2003). While these proteins are adsorbed from fluids such as blood, some may derive directly from the cells themselves, and this protein layer is occasionally referred to as the extra-cellular matrix (ECM). The integrin receptors then become organised into focal adhesion plaques (Burridge et al., 1988), which reorganise the cytoskeleton and lead to cell spreading (Clark and Brugge, 1995). Interestingly, a study by Hu et al. (2001) investigated the basis of phagocyte adhesion to biomaterial surfaces via the Mac-1 integrin cell receptor. This receptor interacted with the P1 D domain of a fibrinogen molecule, but the process of adsorption of the fibrinogen molecule to the material surface denatured the protein and exposed a second potential ligand, P2 (gamma 377395), for Mac-1. These authors suggest that this explains why Mac-1 bearing cells will bind fibrinogen when it is adsorbed to a biomaterial surface, but not when in solution, and thereby initiate inflammation. Alternatively, some surfaces, for instance supported lipid bilayers (SLB), tend to inhibit cell adhesion and it is likely that this happens because of the SLB's ability to resist protein adsorption onto its surface. Andersson et al. (2002) have shown this by comparing protein adsorption on SLBs with that on TiO2 surfaces by quartz

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crystal microbalance dissipation, demonstrating that SLB reduced protein adsorption and epithelial cell adhesion by two orders of magnitude. The time interval between cells attaching to a biomaterial surface and the measurement of adhesion is also important, because the process of full attachment may take minutes, hours or even days. Initial contact between cells and an ECM coated surface may be via very few, possibly even one receptor. Cell shape tends towards the spherical at this stage. However, the cell then undergoes flattening, alignment and spreading, leading to much stronger bonding to the surface, simply because more receptor-ligand bonds are made, probably within a focal adhesion plaque. (For a good review of this process and its implications see Pierres et al. (2002).) These authors interestingly remark upon the two extreme models of the spreading process, one assuming that spreading is a biologically driven process, reliant on rearrangements in the cytoskeleton and localisation of receptors in the adhesion focus, the other suggesting that flattening is due to a balance of adhesive forces with forces tending to resist shape change in the membrane. In any event, assays measuring attachment after just a few minutes will give quite different results from those measuring forces of attachment after some hours. Huang et al. (2003) found that the force needed to detach a chondrocyte from a glass surface by a shear force generated by a vertical cantilever increased from 34 nN to 388 nN monotonically over a period of 6 hrs.

27.3.2 Cell detachment during an assay From the above it is clear that the adhesion of most cells to biomaterial surfaces involves a chain of sites for potential mechanical failure manifest during the deadhesion process. These may be listed as · failure of adhesion of the protein layer to the biomaterial surface · mechanical failure within the protein layer · failure of the cell receptor bond to the protein ligand (this is the failure usually assumed to occur in such studies) · failure of the receptor to remain in the cell membrane · mechanical failure of the cell membrane, either in the bilipid layer, or the cytoskeleton, or detachment of the bilayer from the cytoskeleton (tether formation). As with any chain, failure will occur at whichever of these locations is the weakest. Unfortunately this may mean that some de-adhesion assays are measuring parameters unrelated to the biomaterial properties themselves, such as membrane integrity. However, this does not mean that the only element above worth considering is the adhesion of the protein to the biomaterial surface, since the biomaterial surface properties are known to affect the pattern of protein

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deposition on the material surface (Cacciafesta et al., 2000) and thus exposure of ligands, and so on. Adhesion of proteins to polymers and metals, and mechanical properties of cell membranes are topics which both have a large literature in their own right, and are beyond the scope of this work. From the pioneering work of Rand and Burton (1964) to more recent studies, it is clear that cell membranes display viscoelastic properties, and thus mechanical failure, even at this level, is likely to be complex and time dependent. However, a recent report by Goldstein and DiMilla (2003) observed Swiss 3T3 fibroblasts, after fluorescent staining, during shear induced detachment from fibronectin coated surfaces in a radial flow chamber. The critical finding was that, even over a range of shear stresses, fluorescent `footprints' were seen to remain on the biomaterial surface after cell detachment, indicating that cell membrane, or possibly cytoskeletal, failure can be important steps in cell detachment. This issue is not new, and was probably first articulated by Weiss (1961). A study by Shao and Hochmuth (1999) addressed another link in this chain of adhesion, i.e., the mechanical anchoring strength of receptors in a cell membrane, to the cell cytoskeleton. They studied the anchorage of beta 2 integrins, L-selectin and CD45 in neutrophil membranes by micropipette suction techniques. Forces to extract a likely single receptor ranged from 25 to 130 pN, suggesting this was a feasible mode of de-attachment of a cell from its substratum.

27.4 Sources of variability in adhesion assays If the binding of cells, used essentially as probes, to a biomaterial surface is to be measured, there must be control of both elements of the assay if any degree of standardisation is to be achieved.

27.4.1 The test material surface Many materials used in assays previously have been produced without regard to the reproducibility of their surface properties from batch to batch. Perhaps more worryingly, different materials are sometimes compared for cell adhesiveness, differences being attributed to chemical surface modification when they could equally be due to variation in surface topography, for instance. The manufacture of many polymers and ceramics often leads to non-uniform surfaces which exhibit crystalline `domains', perhaps of varying surface energy. While population type assays may average out such variation over the surface, single cell based assays must be conducted in high numbers of replicates and variability must be expected. Similarly, differing manufacturing processes and later polishing leads to surfaces of varying surface roughness, both at the micron and nanometer level (for a review of methodologies for measuring roughness in this context, see Tomlins et al., 2004).

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It is therefore pertinent to ask if these variations influence the strength of binding of cells to that surface. Richards (1996) compared the attachment of fibroblasts to plastic, titanium and steel samples by measuring the total area of focal adhesion of cells and found that material roughness had no effect on the area of adhesion. Subsequently Lampin et al. (1997) studied adhesion of chick embryo vascular and corneal explants on PMMA for a series of different roughnesses generated by sandblasting. The potential for adhesion increases with increasing roughness but this may reflect a thicker ECM laid down on the rougher surface as revealed by electron microscopy. Conversely, Anselme et al. (2000) studied the adhesion of osteoblasts to a titanium alloy either polished or sandblasted. They found that cells on the rougher sandblasted surfaces never reached confluence and had a stellar shape, and concluded that the organisation of the surface (as measured by a fractal dimension parameter) was the critical factor in adhesion, with lower adhesion being associated with the less organised (rougher) surfaces. However, Delligianni et al. (2001) found that the adhesion of human bone marrow cells to hydroxyapatite surfaces polished with various grades of grit showed a marked increase with increasing roughness. Interestingly, Hallab et al. (2001) found that the increase in adhesion with surface roughness seemed to occur with the polymeric materials they tested, and was not found in metals. Indeed they conclude that surface energy is a more important factor in determining cell adhesion than surface roughness, although it is, of course, difficult to separate the effects of these two variables. Lange et al. (2002) used a model of MG-63 cells adhering to titanium surfaces of varying roughnesses, and made an interesting observation that cells spread more easily on smooth surfaces, and that integrin expression increased on rough surfaces, an observation also noted by Linez-Battaillon et al. (2002). As above, a recent study by Ponsonnet et al. (2003) indicates that surface free energy derived parameters may be more important than simple roughness in determining certain stages, such as spreading, of cell adhesion to biomaterials. Thapa et al. (2003) has taken an approach which states that nano-, rather than micro-scale roughness is of biological importance, and by alkali or acid treatment of test surfaces of PLGA, PCL and PU produced nanoscale surface dimensions. These results suggest that the adhesion of bladder smooth muscle cells is enhanced by nanoscale roughness. These, and several other reports in the literature, suggest a confusing picture of the relationship between material roughness and cell adhesion, but this is likely to be a result of the varying experimental procedures used; metals versus polymers and ceramics, initial adhesion versus spreading; roughness produced by sandblasting or grit polishing, or various ways of actually measuring roughness. It is also likely that the important interaction is between the proteins of the ECM and the surface, before cells actually adhere and spread, which suggests that the surface roughness of the cell itself is of importance.

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27.4.2 Variability in cells used as probes It is obvious that cells of various phenotypes are used in cell adhesion assays, and that these are not necessarily always cells of human, or even primate, origin. Arguments in favour of using different cell types are clear. An implant material intended for use as a bone replacement should be tested against osteoblasts, while one for corneal repair should be tested against corneal epithelial cells. Similarly, an argument can be made for the need for reproducibility over time and between different materials, and between different laboratories and manufacturers. This has probably led to the use of immortal cell lines in such assays, as these grow in an essentially clonal fashion with little change in phenotype with time, and are widely available, but transformed cells may not fully reflect the phenotype relevant to the biomaterial's intended application. This has led to a plethora of different cell types, both primary cultures (freshly isolated and usually of a natural phenotype) and transformed lines being used for cell adhesion studies, with little attention being focused on the effect of this variation. Over a decade ago, Kirkpatrick and Mittermayer (1990) suggested the idea that general screening for biomaterial toxicity should employ easily grown transformed cell lines, and then candidate materials should be screened in an invitro phase using early passage primary cultures. The idea was again addressed by Cenni et al. (1999) in biomaterial toxicity tests. However, some variables in cultures used for testing cell adhesion seem to be regularly overlooked, the major variable being cell senescence. Senescence in primary cultures Isolation of cells from a human donor source, then cultured in flasks may appear to be an attractive source of cells for adhesion assays. Unfortunately, it has long been known that once isolated, such cells have a limited lifespan, the culture reaching a situation defined as senescence after a predetermined number of population doublings (PDs). This limit to proliferation is referred to as the Hayflick limit. This probably does not reflect any fundamental change in the cells after removal from the donor, but more likely a reflection of the acceleration of cell division which occurs in culture. Many cells in vivo, which are not terminally differentiated, probably divide very slowly, perhaps once every few weeks, months or sometimes even years. In culture such cells may divide every two days, thus reaching senescence much faster. Unfortunately, this rate of division is necessary to achieve cell numbers required for many types of adhesion assay prior to frozen storage. To overcome this problem some authors have suggested using early passage cell isolates, and this does, to some extent, overcome the problem. However, it implies that relatively small cell numbers can be generated, especially if the initial biopsy must, of necessity, be small; the corneal epithelium is an example.

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The second problem with this approach is less obvious. A senescing primary culture does not imply that every cell is at an identical stage of senescence. Some cells in the culture will be senescent even in a relatively young culture, and it is the proportion of senescent cells in the culture which increases with increasing PDs. This would affect predominantly single cell type assays of adhesion, when the status of individual cells chosen for assay are unknown. To these authors' knowledge, no comprehensive study of the effects of cell senescence on adhesion to biomaterials has yet been carried out. Overcoming culture senescence There are currently two major commonly used routes to modifying primary cultures to overcome senescence. Firstly, a construct incorporating a so-called anti-oncogene may be introduced into the cell. The concept here is that such anti-oncogenes overcome the effects of genes which act as an `emergency' check on the cell cycle in the event that a mutation in the cell allows escape from cell cycle control. A widely used construct contains the viral gene known as Large-T antigen, which overcomes the effect of the cell cycle `gatekeeper' gene p53. This method has two disadvantages. Firstly, some considerable time may elapse after transfection before clonal outgrowths from culture suggest that some cells have escaped cell cycle control. Secondly, the cell phenotype usually tends to revert to a de-differentiated, more plieotropic appearance. After all, these cells essentially become transformed cell lines. In order to alleviate the problem, some workers have devised constructs with environmentally sensitive Large-T promoters, perhaps temperature or metal ion concentration dependent. The transformed cells can then be grown through the required number of PDs to obtain appropriate cell numbers, and then an adjustment of temperature or medium metal ion concentration is made to turn off the Large-T gene. The cells may then revert to a more organotypic phenotype (for an example see Noble et al., 1993). A second approach to the problem of culture senescence works in cell types in which senescence is related to the shortening of the chromosome telomeres which occurs during chromosome replication. In some cell types (for example T-lymphocytes) activation of the gene coding for the enzyme telomerase allows extension of the telomeres once more, and senescence is delayed. A construct is available which contains the telomerase gene (the so-called tert construct), which can be introduced into cells via a retroviral vector, and appears to delay senescence, possibly indefinitely, in those cell types which are susceptible. This approach usefully appears to retain much of the original cell phenotype, and seems a good candidate for the production of cells for adhesion assays.

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27.5 Methods of assaying cell adhesion in current use No single method appears to have become dominant in the assay of cell adhesion to biomaterials, and different methodologies seem to arise within individual laboratories. Some examples of different methods in the literature are presented below, but these are in no way intended to be exhaustive. A useful review of a selection of these methods has recently been published by Missirlis and Spiliotis (2002).

27.5.1 Cell population methods A number of workers have used a method which involves allowing attachment of cells to a biomaterial surface, and then attempting to detach cells with a jet impingement method. The number of cells remaining can then be measured using microscopy techniques (e.g. Bundy et al., 1994). This represents an attempt to quantify the shear stress needed to erode cells from the material surface. However, a later paper by the same group (Richards et al., 1995) showed that cells were being eroded from the surface chiefly by membrane rupture, suggesting the technique of jet impingement may not have been suitable for quantification of the force of adhesion of cells to biomaterials. Parallel-plate flow chambers have been used extensively to attempt to measure adhesion of cells to biomaterials. This is a valid approach for a straightforward assay provided the shear stress acting on cells can be calculated (Brooks and Tozeren, 1996). The number of cells left attached to biomaterial surfaces in such simple assays as flow chambers, (or even simply exposing biomaterials to cell suspensions) can be problematic. If simple phase contrast microscopy is not possible, perhaps because of the opaqueness of the biomaterial, cells may be detected by colorimetric means (Humphries, 2001) by indirect immunofluorescence of adhesion plaque components with image analysis (Hunter et al., 1995) and using radioactively labelled cells (Nayab et al., 2003). A useful approach has been developed by Dexter et al. (2003), who developed a bioluminescent-based ATP assay which is capable of enumerating bacterial and mammalian cells differentially, an important capacity if biomaterials are to be developed which bind human, but not bacterial cells. Goldstein and DiMilla (1997, 2002) used a radial flow chamber to examine the adhesion of cells to substrates. Here cells were introduced into a chamber and allowed to settle quiescently for a few minutes so as to allow attachment but not spreading. Flow was then applied in a divergent radial pattern over the surface such that the shear stress at any point along the chamber radius could be calculated. This technique does not seem to have been widely adopted by other authors. In an attempt to generate a simple and reproducible adhesion assay with high throughput, Reyes and Garcia (2003) used a centrifugation based assay which

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they claim generates robust statistics for adhesion analysis. They simply seeded cells onto substrates in a 96-well plate and after rinsing, centrifuged after inversion, and used a fluorescent technique to count the remaining cells. This is an interesting technique, and unusual in applying a normal, rather than a shear force to detach the adherent cells. A novel approach has been taken by Nordon and colleagues (2004) in immobilising a binding ligand (protein A) in hollow fibres and then allowing attachment of cells to the modified surface. Adhesion was quantified by calculating the flow rate of fluid through the hollow fibre, thus yielding a shear stress value at the fibre wall, and counting the number of cells detaching. However, the standard flow cell is still a parallel sided chamber with an inflow and outflow creating a laminar flow pattern over adherent cells on the lower surface of the chamber. The height of the flow cell has to remain small to allow visualisation of the base by microscopy, but also has to be larger than the cell diameter by two orders of magnitude to ensure that wall stress can equal shear stress on the cells. An interesting observation has been made by Boyle (2004) when observing the attachment of bacterial cells to plastic surfaces in a square section flow cell using videomicroscopy over long time periods (often many hours). It is often assumed that the increase in cell numbers on the surface is due to the accumulation of cells over the whole course of the experiment. However, Boyle showed in this case that cells were adhering and then de-adhering over the timecourse of the experiment. However, more cells were attaching as the time of the experiment increased, giving the appearance of steady accumulation. This type of analysis should be done for mammalian cells as it would have important consequences for adhesion assays of the flow cell type.

27.5.2 Single cell methods Methods for measuring the adhesion of single cells generally fall into one of three categories, namely microcantilever displacement methods, optical methods and micropipette methods. Micropipette methods have the longest history of these three, being originally employed by Mitchison and Swann in 1954. This essentially simple but technically demanding technique uses a drawn glass pipette with a tip diameter of less than ten microns, filled with buffer fed from a movable reservoir such that the pressure in the pipette can be finely adjusted by vertical movement of the reservoir. The micropipette is then loaded into a micromanipulator rig and observed under a microscope, the system being similar to that traditionally used by cell electrophysiologists. The micropipette can then be used to remove a cell from a surface, perhaps of a bead or planar substrate, and the force required calculated. Originally used for the measurement of the bending moment of red blood cell membranes (Rand and Burton, 1964), this method poses some difficulties in this application since the micropipette tip

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diameter has to be similar to the cell diameter to reduce energy loss due to cell deformation during the experiment. Nevertheless, values for the removal of red cells from surfaces have been reported using a modification of this method (Bowers et al., 1989). More recently, the manipulation of cells close to a surface using optical tweezers has become a possibility. This modification of a technique, originally developed to trap virus particles, relies on a phenomenon whereby a particle with a refractive index greater than its surrounding medium, and a diameter which is much less than the value 2= (where  is the wave number of the light from a monochromatic laser focused to the spot at which the cell is located), experiences an optical force which tends to move it towards the centre of that spot. The force is composed of two elements, one which tends to drive the particle up the gradient of light intensity (towards the centre of the focused spot) and a second due to light scattering which tends to drive the particle along the beam axis. Clearly the axial intensity gradient must exceed the displacement force generated by scattering, so that the cell becomes trapped in the focused laser spot and can be manipulated. To achieve this, a wide beam needs to be focused in a microscope objective in order to generate a small spot such that the optical gradient at the spot is large. This has been used to manipulate fibroblasts near glass surfaces, but as yet cannot generate enough force to detach a cell from a substrate, though it can form tethers. This approach has been used by Huang et al. (2003) to investigate tether formation in chondrocytes by attaching microbeads to a cell surface and then trapping the beads in the tweezers. By pulling the bead away from the cell, a tether is formed, and the force required can be measured. The most widely used technique for manipulation of single cells adhering to a surface is to employ some form of microcantilever. This is used in one of two configurations, either in the traditional atomic force microscope (AFM) mode which generates, or measures, a normal force, or in shear mode. In either approach, the deflection of a microcantilever of known stiffness is monitored by the movement of a laser beam reflecting off the cantilever upper surface. The cantilever is usually fitted with a fine tip which contacts the cells. As an example of shear measurement, Yamamoto et al. (1998) arranged for a microcantilever to have a right-angle crank in the beam so that the tip could be brought up to a cell adhering to a glass surface, and then removed in shear mode. The bending of the vertical part of the beam allowed measurement of shear force to be made. By using a CCD based microscopy system, the adhesion area between cell and substrate could also be measured. They succeeded in measuring detachment forces per unit area of attachment of the order of 530 to 750 Pa. It is noteworthy, in the light of comments in the section above, that these workers found that cells prepared on different days, but by exactly the same methods, yielded variable results, whereas those prepared on the same day gave reproducible results between replicates. Huang and co-workers (2003) used

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a similar system with a vertical beam (a cytodetacher) to measure the forces of adhesion of chondrocytes to a glass surface over time. As mentioned previously, Razatos and co-workers (1998a) used a standard AFM configuration, and suggested a micron size bead of a putative biomaterial attached to the AFM tip. This could then be used to measure the attraction/repulsion of a material from a bacterial lawn. They also inverted the geometry, by growing bacteria on an AFM tip. A similar approach was adopted by Thie et al. (1998) to investigate interactions between JAR cells and uterine epithelial monolayers. They glued a single Sephacryl S-1000 bead of 80 microns diameter to the AFM cantilever tip, and then grew the JAR cells on this surface, providing a functionalised tip for probing epithelial monolayers in a so-called wet cell, in cell culture medium. An interesting but completely different approach was used by Patel et al. (2000) where a bioadhesive polymer was exposed to buccal epithelial cells from suspension. The AFM was then used to image the cell surface at very high resolution in order to identify the macromolecules of the polymer adsorbed to the cell surface.

27.6 Conclusion The attachment of cells to biomaterial surfaces is of considerable importance in the success of many implanted devices. However, techniques employed to measure this adhesion are varied, and probably determine a broad range of parameters, each related to adhesion, but not necessarily the fundamental factor. Indeed, it may well be that the very term `cell adhesion' is misleading in that it implies one parameter, when it is composed of a series of interactions which add up to the overall effect of a cell adhering to a biomaterial. Similarly, assays of cell adhesion often have varying objectives, sometimes the investigator is interested in measuring the force of adhesion between a single cell, or even a single receptor, and a biomaterial, sometimes they are interested only if one biomaterial appears to prove more attractive to a cell than another. Reproducibility of the surface to be tested has received attention in the literature with investigations into the effect of surface roughness being numerous. However, it may now be important to turn attention from sculpturing at the micron range to the nanometer range. Rather limited attention has been given to the problem of choosing cells to be used in adhesion assays, with expediency sometimes overcoming the need for reproducibility. Indeed, perhaps some of the variation in results seen between various workers is due to wide variation, not just of cell types used, but of the manner in which the cells are cultured and the environment in which they are tested. The geometries used in adhesion studies centre on a small number of basic approaches to applying forces to cells, viz. flow chambers of various designs (also spinning disks); simple rinsing; jet impingement; centrifugation; micropipette techniques; microcantilever and AFM techniques; optical tweezer techniques.

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It is interesting that in 1997, Otto et al. commented on the importance of international standards for biomaterial interactions with cells, using a standard testing panel. It seems clear that to achieve this for cell adhesion measurement, more than one technique will be necessary, probably with more than one cell preparation employed. In any event, assays of cell adhesion are likely to remain of real practical importance so long as a theoretical prediction of cell adhesion remains elusive (for a discussion see Vitte et al., 2004).

27.7 References Acaturk TO, Peel MM, Petrosko P, LaFramboise W, Johnson PC, DiMilla PA (1999) J.Biomed Mater. Res. 44 (4) 355±370. Andersson A-S, Glasmastar K, Sutherland D, Lidberg U, Kasemo B (2002) J. Biomed. Mater. Res. 64A 622±629. Anselme K, Bigerelle M, Noel B, Dufresne E, Judas D, Iost A, Hardouin P (2000) J. Biomed. Mater. Res. 49 (2) 155±166. Bell GI (1978) Science 200 618±627. Bowers VM, Fisher LR, Frances GW, Williams K (1989) J. Biomed. Mater. Res. 23 1453. Boyle JD (2004) Pers Comm. and PhD thesis, Dept of Engineering, University of Exeter, UK. Brooks SB, Tozeren A (1996) Computers and Fluids 25 (8) 741±757. Bundy KJ, Roberts OC, O'Conner K, McLeod V, Rahn B (1994) J. Mater. Sci. Mater in Med. 6 500±502. Burridge K, Faith K, Kelly T, Nuckolls G, Turner C (1988) Ann. Rev. Cell Biol. 4 487± 525. Cacciafesta P, Humphris A, Jandt K, Miles M (2000) Langmuir 16 8167±8175. Cenni E, Ciapetti G, Granchi D, Arciola C, Savarino L, Stea S, Montanero L, Pizzoferato A (1999) Toxiclogy in Vitro 13 801±810. Charm S, Kurland G (1974) in Blood Flow and Microcirculation. J. Wiley and Sons NY, 72. Chen S, Springer TA (2001) PNAS 98 950±955. Cheng Z (2000) J. Biomechanics. 33 23±33. Cheng Z, Bao G, Wang N (2000) Ann. Rev. Biomed. Eng. 2 189±226. Clark EA, Brugge JS (1995) Science 268 233±239. Costerton JW, Lewandowski Z, Caldwell DE, Korber DR, Lampin-Scott HM (1995) Ann. Rev. Microbiol. 49 711±745. Decave E, Garrivier D, Brechet Y, Fourcade B, Bruckert F (2002) Biophys. J. 82 2383± 2395. Deligianni D, Katsala N, Koutsoukos P, Missirlis YF (2001) Biomaterials 22 87±96. Dexter SJ, Camara M, Davise M, Shakesheff K (2003) Biomaterials 24 27±34. Garrivier D, Decave E, Brechet Y, Fourcade B (2002) Eur. Physicsl J. E. 8 79±97. Goldman AJ, Cox RG, Brenner H (1967) Chem Eng Sci. 22 637±660. Goldstein AS, DiMilla PA (1997) Biotech and Bioeng. 55 616±629. Goldstein AS, DiMilla PA (2002) J. Biomed. Mater. Res. 59 665±675. Goldstein AS, DiMilla PA (2003) J. Biomed. Mater. Res. A. 67A 658±666. Gosiewska A, Rezania A, Dhanaraj S, Vyakarnam M, Zhou J, Byrtis D, Brown L, Kong W, Zimmerman M, Geesin JC (2001) Tissue Eng. 7 267±277. Hallab NJ, Bundy KJ, O'Connor K, Clark R and Moses R. (1995) J Long-term Effects.

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Med. Implants. 5 209±231. Hallab NJ, Bundy KJ, O'Connor K, Moses R. and Jacobs J (2001) Tissue Eng. 7 55±71. Hersel U, Dahmen C, Kessler H (2003) Biomater. 24 4385±4415. Hochmuth RM, Mohandas N (1972) J. Biomech. 5 501. Hu WJ, Eaton JW, Tang L (2001) Blood 98 1231±1238. Huang W, Anvari B, Torres J, LeBaron R, Athansiou K (2003) J. Orthopaed. Res. 21 88±95. Humphries MJ (2001) Molec. Biotech. 18 57. Hunt JA, McLaughlin PJ, Flanagan BF (1997) Biomater. 81 1449±1459. Hunter A, Archer CW, Walker PS, Blunn GW (1995) Biomater. 16 287±295. Kirkpatrick CJ, Mittermayer C (1990) J. Mater. Sci. Mater in Med. 1 9±13. Kochen J (1966) Proc 1st Int Conf Haemorheol. Iceland. 1 455. Lampin M, Warocquier-Clerout R, Legris C, Degrange M, Sigot-Luizard M (1997) J. Biomed Mater. Res 36 99±108. Lange R, Luthen F, Beck U, Rychly J, Baumann A, Nebe B (2002) Biomolec Eng. 19 255±261. Linez-Batallion P, Monchau F, Bigerelle M, Hildebrand H (2002) Biomolec. Eng. 19 133±141. Missirlis YF, Spiliotis AD (2002) Biomolec Eng 19 287±294. Mitchison GP, Swann MM (1954) J. Exp. Biol. 31 443. Nayab S, Shinawi L, Hobkirch J, Tate TJ (2003) J. Mater. Sci. Mater in Med. 14 991±997. Noble M, Groves AK, Ataliotis P, Morgan J, Peckham M, Partridge T, Jat PS (1993) Neuroprotocols: A companion to methods in neurosciences. 3 189±199. Nordon R, Shu A, Camacho F, Milthorpe B (2004) Cytometry A 57A 39±44. Otto M, Klein C, Kohler H, Wagner M, Rohrig O, Kirkpatrick CJ (1997) J. Mater. Sci. Mater in Med. 8 119±129. Patel D, Smith J, Smith A, Grist N, Barnett P, Smart J (2000) Int. J. Pharm. 200 271±7. Pierres A, Benoliel A, Bongrand P. (2002) Eur Cells and Maters. 3 31±45. Ponsonnet L, Reybier K, Jaffrezic N, Compte V, Lagneau, Lissac M, Martelet C (2003) Mater. Sci. and Eng. C 23 551±560. Rand RP, Burton AC (1964) Biophys J. 4 115. Razatos A, Ong YL, Sharma MM, Georgiou G (1998a) J Biomater. Sci. Polmer Ed. 9 1361±1373. Razatos A, Ong YL, Sharma MM, Georgiou G (1998b) PNAS 95 11059±11064. Reyes CD, Garcia AJ. (2003) J. Biomed Mater. Res. 67 328±333. Richards R (1996) Injury. Int J. of the Care of the Injured 27 38±43. Richards R, apGwynn I, Bundy KJ, Rahn BA (1995) Cell Biol Int. 19 1015±1024. Shao J, Hochmuth RM (1999) Biophys J. 77 587±596. Templeman L, Park S, Hammer DA (1994) Biotechnol Prog. 10 97±108. Thapa A, Webster TJ, Haberstroh K (2003) J. Biomed Mater. Res. A. 67A 1374±1383. Thie M, Rospel R, Dettmann W, Benoit M, Ludwig M, Gaub H, Denker HW (1998) Human Reprod. 13 3211±3219. Tomlins P, Michalovsky S, Vadgama P, James SL (2004) NPL A report, in press. Vitte J, Benoliel AM, Pierres A, Bongrand P (2004) European Cells and Materials 7 52±63. Vroman L, Adams A, Klings M (1971) Fed Proc 30 1494. Weiss L (1961) Exp. Cell. Res. 8 141±153. Wilkins K, Hanlon G, Martin G, Marriott C (1990) Int. J. Pharmaceut. 58 165±174. Yamamoto A Mishima A, Maruyama N, Sumita M (1998) Biomaterials 19 871±879.

28

Protein adsorption to surfaces and interfaces

B M I L T H O R P E , University of New South Wales, Australia

28.1 Introduction Surfaces are often represented as flat or smooth planes on which molecules settle and adsorb. This is far from the case for current and future biomaterials. It is, however, a useful starting point in order to examine the various types of biomaterial surfaces and their important interactions from a protein macromolecular and/or biochemical point of view. The development of an understanding can progress from this point to include the effects of the various properties of the surface(s) and protein(s) under examination. The concepts of a surface and an interface also deserve careful examination at the molecular level as the junction between two separate environments may also represent a series of complex interactions and topologies. The junction may be `obvious' at a macroscopic level, yet impossible to define accurately at the molecular level with graded, deep surfaces. The key to understanding protein adsorption is the understanding of the adsorption processes involved, the driving forces for the processes, the free energy of adsorption, and the nature of the surface or interface involved, the competing processes and any synergistic processes. The main distinction to draw when first approaching the problems of protein adsorption to surfaces is to distinguish `non-specific' from `specific' adsorption. Specific adsorption (specific binding) usually involves a high affinity binding to a very defined chemistry, such as may occur with a lectin to a surface bound sugar. Everything else is non-specific, even though the process may be reversible. Free energies are usually associated with equilibrium situations and are most easily estimated at equilibrium. The various components that make up the free energy of adsorption will be considered below. However, much work is still required in this area, especially for non-classical surfaces or interfaces. Many protein adsorption studies have been undertaken after sufficient time that equilibrium may be assumed. However, protein adsorption at the biomaterial interface may not ever reach equilibrium before key biological events occur, so we will need to develop tools that estimate behaviour far from

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equilibrium as well. This aspect also is not well developed from a theoretical point of view. Protein deposition, often studied in isolation, does not occur in isolation in vivo and the effects of water, lipids (Iwasaki et al., 1999) and other biochemical molecules are also important on the processes driving protein adsorption and, perhaps, degradation. Thus the system in which protein adsorption is being studied is another important criterion.

28.2 Classification of biomaterials surfaces and interfaces As far as biomaterial interactions with proteins are concerned, features on the Angstrom and nanometer scale are important. Features may be physical (e.g. shape, distribution of domains), chemical or electromagnetic. The features essentially define the surface, or the interface that occurs between two distinguishable phases. In general, a `surface' occurs where there is a step function in properties between the phases that defines the position in space of the interface. This may be a change in state, i.e., solid-liquid, solid-gas or liquidgas, or the formation of an immiscible boundary, e.g., oil and water. We may classify surfaces into a wide variety of different classes depending on which feature set is used first. From the historical perspective, and from the development of theory, a physicochemical feature set is most commonly used first. This is the molecularly `hard' surface. Often treated as molecularly flat as well, this surface does not allow penetration of protein or other molecules and is chemically stable. The surface itself may have larger scale roughness and there are now many studies on the effects of nanoscale features of these `hard' surfaces. On the other end of the scale are the thick hydrogel interfaces that allow intermolecular penetration of large macromolecules ± reminiscent of the glycocalyx surrounding many cell plasma membranes or the hyaline cartilage surface. In between are surfaces that do not allow large molecules to penetrate, but do re-arrange (e.g. many polyurethanes), surfaces that allow small molecules to partially penetrate, surfaces that are nanoporous or microporous, surfaces that are hydrogels and surfaces that are mixtures of different characteristics on the nano or micro level. At yet another level are the graded interfaces, where the two different phases can be distinguished at a macro level, but the change in properties between the distinguishable phases occurs as smooth continuous functions rather than discontinuous functions (e.g. interpenetrating polymers). There are also the surfaces that are environmentally `switchable' like polyNIPAAm (poly Nisopropylacrylamide). These are very useful materials, but very hard to treat theoretically to determine protein adsorption characteristics. Biological surfaces and interfaces form a fascinating extra set of classes of surfaces and interfaces. In many of these, water forms a continuous underlying

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connection, but the polymeric components are spatially separated (e.g. cartilage in joints). Cell membranes appear to be easily identifiable spatially, but the various components of the membrane do not lend themselves to simple spatial definition at the Angstrom level. Tissues, cells and organelles, have distinguishable phases, but the interfaces between the components are also difficult to define. For the biomaterials scientist or engineer, much of what occurs at those interfaces is of interest, as it helps to define the types of interactions that are desirable for the biomaterials of the future.

28.3 Non-specific adsorption to hard surfaces We begin with non-specific adsorption to hard surfaces as these constitute the vast majority of situations to date. Most surfaces of biomaterials, to date, are of the classical solid type. These include ceramics (e.g. Al2O3, ZrO2, Ca10(PO4)6(OH)2), metals (cobalt chrome, titanium alloys, stainless steel) and many polymers (e.g. polyethylene, polypropylene, polystyrene, polycarbonate, polyethyleneterephthalate, polysulphone, polyethyleneetherketone, polytetrafluoroethylene, polysiloxanes). These surfaces are impervious to most molecules (ignoring, for the moment, solvation in organic fluids, and plasticisation of polymers). Not only that, but the surface inhomogeneities are such that the materials can be treated as homogeneous for both large and small molecules alike. The main, commonly used, measure of surface chemistry is its free surface energy (also known as the surface tension for liquids). The total free surface energy is normally estimated by measuring the contact angle a (small) drop of water makes with the solid surface. This measures the effective force generated by the various interfaces (air-water, air-material, water-material) at the junction of all three. The Young equation relates the surface free energies of the solidvapour interface ( SV ), the solid-liquid interface ( SL ) and the liquid-vapour interface ( LV ) to the contact angle ().

SV ÿ SL ˆ LV cos 

28:1

Values of LV are available from measurement of liquid drop pressures. For pure liquids on solids, SL will represent the total energy of interaction of the liquid with the solid, whilst SV is the effective energy required in order to generate a new surface on the solid in air. The relationship between SL and S (the solid surface free energy) is usually assumed to be given by Fowkes (1965), Owens and Wendt (1969), Kaelble (1971) and Schakenraad et al. (1986):

SL ˆ S ‡ L ÿ 2… Sd Ld †1=2 ÿ 2… Sp Lp †1=2

Ad

Ap

28:2

The quantities and refer to the dispersive and polar components of the surface free energy for component A. The validity of this formula has been questioned for the summation of polar forces (there are various models) (Wu, 1971, 1980; Fox and Zisman, 1950, 1952a,b; Zisman, 1964; van Oss, 2002).

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The quantity S , the free energy of forming the surface in vacuo, is the value normally quoted for solid materials, although Vogler (1998) argues strongly for 0 the use of `Water Adhesion Tension',  0 ˆ LV cos , a close relation to the 0 …1 ‡ cos †. Young-DupreÁ equation for the work of adhesion, Wa ˆ LV It may be argued that either measure is equally valid for the purposes of comparison of protein binding behaviour in the region of values of S or  0 normally associated with biomaterials. I will tend to use the S values due to more common usage in the literature. The measure one chooses to use is of concern when investigating the mechanisms and kinetics of binding. As a general rule, surfaces with a low surface energy ( S < 30 mJ mÿ2) are hydrophobic and those with a high surface energy ( S > 73 mJmÿ2) are extremely hydrophilic (wetting). The range of intermediate energies (40 to 70 mJ mÿ2) is one of considerable interest as most biomaterials fall into this region. Most biological surfaces and interfaces, unless in contact with air, tend to be in the range >73 mJ mÿ2 as they are covered with highly hydrophilic moieties (-OH, -COOÿ, -NH3+, -PO4Hÿ, -SO3ÿ and others).

28.3.1 Adsorption isotherms When discussing protein adsorption it is essential to define the exact question that is being asked. Much confusion is possible if the various states are not well understood and the terminology defined. The classical situation, which will be assumed here, is a solution of a protein or proteins in a buffered aqueous solvent in direct contact with a material surface. In this classical situation there will be protein that is in direct contact with the surface, which is called adsorbed protein here, and there may be further layers of associated protein (De Cupere et al., 2003; Pernodet et al., 2003; Santos et al., 2002; Krishnan et al., 2003; Voros, 2004). There is the equilibrium state and the kinetics of adsorption potentially to consider. The equilibrium state is often described by an adsorption isotherm, usually a Langmuir (Gould, 1968) or Freundlich isotherm although there are many others depending on the adsorption model assumed (Linden et al., 2002; Lucassen-Reynders et al., 2004; Bosma and Wesselingh, 2004; Dias-Cabral et al., 2003; Bentaleb et al., 1998; Wassell and Embery, 1996). Langmuir:

q…C† ˆ

qmax C C‡K

where q…C† is the amount of adsorbed protein, C is the concentration of protein in the bulk solution and K is a fitting constant. Freundlich:

m ˆ ab ;

where m is the mass adsorbed per unit area, a is a constant related to the maximum mass adsorbed,  is the mole fraction of protein in bulk solution

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…0    1† and b is a constant. For monlayer adsorptions b can be related to the Gibbs free energy of adsorption by h
Gads i ˆ ÿ…1=b†RT, where h
Gads i is the molar free energy of adsorption, R is the Rydberg gas constant and T the temperature in ëK (Martinez et al., 2000). In general, partly due to ease of measurement, most studies of protein binding look at the adsorbed protein left after a series of rinses. This remaining protein is often referred to as `irreversibly bound' protein. When irreversibly bound protein is analysed as a function of the bulk concentration of originally overlying protein solution, it often conforms to an adsorption isotherm of the Langmuir or Freundlich type. Care must be exercised in interpreting or using the apparent constants that come from these studies as they relate to a non-equilibrium condition (rinsed).

28.3.2 Kinetics of adsorption The kinetics of adsorption relate strongly to the design of the experimental setup, as well as the surface, protein and protein concentrations, and buffer variables (salts, salt concentration, pH and temperature). There are many kinetic studies of protein adsorption (see for example, Kasemo, 1998; Al-Malah and Mousa, 2002; Ramsden, 1993; for more general articles: Krishnan et al., 2004; Lee et al., 2004; Dupont-Gillain et al., 2003; Fong et al., 2002 for specific examples). Rates of adsorption and formation of multilayers, are heavily dependent on the individual experimental conditions. Conditions of temperature, pH, ionic strength and salts all affect the outcome. In general, though, kinetic experiments may be used to distinguish models of adsorption (e.g. reversible adsorption, reversible with some conversion to irreversible adsorption, multilayer adsorption) as well as giving estimates of rate constants. Low concentrations and stationary fluid conditions may lead to diffusion limited kinetics. The effective diffusion velocity for proteins is of the order of 10 m/s at 310 ëK so that, after a few seconds of adsorption the layer of solution closest to the surface becomes relatively depleted in protein. As the bulk protein concentration increases, this effect is reduced. For very dilute protein solutions, very long times may be necessary to establish equilibrium. Well-mixed fluid over the surface will reduce the diffusion limitation, but high surface fluid shear can affect the formation of associated adsorbed protein layers. In order to probe the kinetics of the adsorbing protein or the equilibrium isotherm of total adsorbed protein, it is necessary to determine the mass of protein in the adsorbed layer (or Gibbs phase; Toth, 2002) without disturbing the system. There are a number of methodologies that provide estimates of these data, providing certain assumptions are fulfilled. Ellipsometry (Elwing, 1999) uses the change in refractive index of the adsorbed layer. The quartz crystal microbalance with dispersion (QCM-D) uses changes in the mechanical loading of a resonating quartz crystal (HoÈoÈk et al., 1998). The resonant frequency and its

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overtones can be shown to be related to the mass adsorbed to the surface, including bound water, whilst the damping of the resonance (dispersion) is related to the viscous damping of the layer. With adsorbed proteins, a model relating mass and viscosity needs to be assumed to derive the mass and viscous constant results from the frequency and dispersion data. Another alternative is based on surface plasmon resonance and/or evanescent wave technology (Stenberg et al., 1991; Freeman et al., 2004; Santos et al., 2004). These technologies use the interaction of the evanescent wave with the adsorbing material, effectively detecting the change in refractive index as protein or other molecules adsorb or desorb.

28.3.3 Surface inhomogeneity Most hard surfaces are not homogeneous at a molecular level. Unless the surface is an orientated element (e.g. gold on a particular lattice orientation) then differing parts of the surface contribute different areas of potential binding energy (dispersive hydrophobic regions, permanent dipoles, hydrogen bonds and charged groups). Random arrival of proteins at these sites in random orientations means that different interactions will occur for protein molecules arriving at different positions on the surface. Also the adsorbed proteins will not readily diffuse, so that as protein molecules (or other molecules) bind to the surface, free areas may be left that are not large enough to accommodate a protein molecule in a way that allows adsorption. The net effect is that the maximum surface coverage is less than the theoretical maximum calculated from protein dimensions and surface dimensions. Surface rearrangement is another potential confounding process. This is especially likely for non-crystalline polymers containing polar groups (e.g. polyurethanes, polyesters, polyamides, polysaccharides). The polymer chains are partially mobile and may rearrange to provide the most hydrophobic surface to air and the most hydrophilic surface to water. These rearrangements take time, so that protein adsorption to a freshly wetted surface may well differ from the same material that has been in contact with water for some time. Protein adsorption may also alter the chain conformation at the surface to provide increased binding energy. It is difficult to distinguish the surface rearranging from the protein rearranging, although a clue to surface mobility is long-term changes in advancing contact angle between surfaces kept wet and surfaces kept dry.

28.4 General rules of non-specific adsorption to flat surfaces Most protein adsorption to solid surfaces is essentially by non-specific binding. There have been attempts made in recent times to cause proteins to bind to specific spots and in specific orientations either for biosensors (Bucur et al., 2004; Jun et

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al., 2004; Zhang et al., 2004) or for cell attachment (Dewez et al., 1998; Goessl et al., 2001; Kasemo, 2002; Luk et al., 2000; Scotchford et al., 2002, 2003; Sugawara and Matsuda, 1996). It is generally agreed that the energetics of nonspecific binding is largely dominated by the dispersive forces. Thus, the generally found phenomenon of reduced total protein adsorption with increased hydrophilicity/surface energy. It must be noted, however, that more `specific' binding is still possible for some proteins with very hydrophilic surfaces. For proteins binding to biomaterials, the driving forces are the interaction energy of the surface with water (the energy measured from water contact angles), the interaction energies of the protein with the surface (more later) and the interaction energies of the protein with water. Interaction energies may be either enthalpic or entropic (or both). The enthalpic energies are associated with the dispersive interactions (often called hydrophobic interactions: Lifschitz, van der Waals, London), and the polar interactions (charged group attractions or repulsions, salt bridges, dipoles, and hydrogen bonds). Rarely, there may also be chemical reactivity resulting in covalent bond formation. Entropic energies are largely dominated by water effects from both surface and protein. Adsorption of the protein occurs when the free energy of the adsorbed system (protein adsorbed to surface) is lower than the solution system (surface in contact with aqueous solution and protein in solution). For small molecules adsorbing to a surface in vacuo it is possible to calculate, a priori, the relevant interaction energies. From the interaction energies, the adsorption isotherm models (Langmuir, Freundlich, Fowler-Guggenheim, Volmer, de Boer-Hobson, ToÂth) are well documented and may be used for some proteins in solution in some simple cases (single small protein in water or simple salt solution). The main requirements for standard adsorption isotherms to be used are small, hard proteins that can move (diffuse) readily across the surface, and a surface that is chemically homogeneous with respect to the size of the protein. General protein adsorption (or adsorption of any large molecule) is a different matter. Protein adsorption to many of the hard surfaces tends to be by hydrophobic interactions of the more hydrophobic parts of proteins with the surface. This results in a change in protein configuration, which may be irreversible. The binding of albumin (Puskas et al., 2004; Song and Forciniti, 2000; Yoon et al., 1999) to polystyrene is reported to be largely by hydrophobic interactions and it is likely that other proteins also show some hydrophobic bonding character. The degree of hydrophobic interaction available from the surface determines the strength of the bond and the degree of conformational change. For example, albumin adsorbs strongly to unmodified polystyrene (microbiological class polystyrene), which makes it a good `blocking agent' for ELISAs, preventing other proteins and molecules from non-specifically binding to the surface. Tissue culture (TC) polystyrene has been modified to provide a certain density of hydrophilic groups on the surface. This increases the water

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binding energy to the surface (higher surface energy) and reduces the free energy of adsorption for albumin. Consequently other proteins are more able to compete for the surface (see Vroman effect, section 28.4.1). These include fibronectin and vitronectin, which are essential for various cells to adhere and spread on the TC polystyrene surface. As the surface energy increases further, the energy available from hydrophobic interactions decreases. This decrease can be attributed largely to the enthalpic and entropic energy costs of removing the bound water from the surface. Under these conditions one would expect only specific interactions of proteins with surface moieties to drive protein adsorption. Protein orientation also has a bearing on this as proteins have very specific configurations for residues capable of forming bonds with surface groups (the polar forces). Thus a protein molecule arriving in one orientation may not bind to a surface, whereas it may bind very strongly in another orientation. Each hydrogen bond adds about 2 to 19 kJ/mole to the interaction energy assuming similar energies to protein folding. Thus the net formation of four or five hydrogen bonds (~50 kJ/mole) will render a protein almost completely `irreversibly' adsorbed (apparent Keq (adsorption) ˆ 109). However, for most hydrogen bonding situations, the net energy of bonding after displacing water is low or even negative. Thus hydrogen bonds alone, unless of a type that provides substantial net energy on formation, will not provide a great deal of energy for driving protein adsorption. The net effect of random orientation of proteins and inhomogeneous surfaces is to provide a complex adsorption process, that can be approximated by ensemble averaged adsorption coefficients and reduced maximum coverage of `irreversibly' adsorbed protein. Over very long periods the adsorbed protein will be observed to slowly change towards the most energetically favoured conformation(s) bound to the surface, but adjacent `binding sites' may have different conformations that are most energetically favoured. Protein orientation and conformation re-arrangement has been observed for several proteins (Agnihotri and Siedlecki, 2004; Wagner and Castner, 2004; Jung S-Y et al., 2003; Ta and McDermott, 2000; Brusatori and Van Tassel, 1999), but is especially well documented for the serum albumins (Carter and Ho, 1994). Serum albumin is found in most animals, and has a reasonably conserved structure amongst the mammals. There are nine loops forming three homologous domains. Overall, the tertiary structure is `heart shaped', close to being an equilateral triangle with 8 nm sides and a thickness of 3 nm. Considerable conformational change can be induced in serum albumin by changes in pH and by binding fatty acids. Although serum albumin is negatively charged and highly water soluble at neutral pH, its fat binding capability is substantial and thought to be the reason for its ability to bind to hydrophobic surfaces. Also, the ability to change conformation to increase fat binding capacity could well be the mechanism for the formation of `irreversibly bound' albumin on hydrophobic surfaces.

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Jung et al. (2003) have demonstrated that fibronectin binds in two preferred conformations to silica at pH 8.0. One is relatively easily displaced whilst the other, minority, conformation is only slowly displaceable. By cycling the pH to 3.2 and back, the majority of the fibronectin is found in a very strongly bound conformation. They postulate that the `weakly' bound form is the form most likely to be found at neutral pH and therefore to be displaceable by larger proteins (Vroman effect). The conformational change is thought to be related to the C domains on fibronectin `shielding' the D and E regions from binding to the surface. At the same time, the interactions are likely to shift from polar to dispersive. This agrees well with the results of Weber et al. (2004) who show, for specific polymer families, a decrease in fibronectin binding with increasing hydrophobicity of polymer surface.

28.4.1 Dynamic binding ± Vroman effect In most applications of biomaterials, the solution first in contact with the biomaterial is a complex mixture of proteins, lipids and salts (usually plasma and lymphatic fluid, but may also be tears, mucus, saliva or other fluid). A few species of protein are normally present in abundance. These species dominate early binding behaviour to the surface as they arrive in abundance. As time progresses, other species of protein that have much lower concentration, but greater effective adhesion will start to appear in larger quantities. This process is not an equilibrium process, but a dynamic, competitive one. It is the underlying basis of the Vroman effect (Krishnan et al., 2004; Jung et al., 2003; Vroman and Adams, 1969; Wojciechowski et al., 1986; Slack and Horbett 1992; Dejardin et al., 1995). The molecular mechanisms underlying the Vroman effect are still subject to debate (Krishnan et al., 2004; Jung et al., 2003). Protein that is apparently irreversibly bound can still be displaced by the (larger) protein. However, the contradictions are really a function of the assumptions made when developing the relevant models to attempt to describe protein adsorption behaviour. If a protein is expected to adsorb and desorb as a single unit, then the irreversibly bound protein should not be displaced. This is the standard model used for kinetic descriptions of adsorption and it works well for small molecules and small, relatively rigid proteins. Large proteins, say, from albumin and larger, have several domains that interact relatively independently with the surface. If the individual domains are modelled as being able to adsorb and desorb independently, but the protein is a linked unit, then the whole protein will tend to remain adsorbed, and probably will not be displaced easily by more of the same protein. However, another protein, whose domain(s) are more favourably bound, or with more domains, will bind to areas of surface that are vacated by the first protein's domain(s). As the second protein will have a lower probability of desorbing before another

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domain of the first protein desorbs, the second protein displaces the first protein over a period of time. The effective `stickiness' of a protein (i.e. the ability to compete for a nonspecific binding site) may be related to the product of the bulk concentration of the protein (Cp) and the effective adsorption equilibrium binding constant (Kadseff). The higher the product, the more effectively the protein will compete for sites. Proteins that are truly irreversibly adsorbed have a Kads-eff of infinity. The rate at which the surface is populated, or the competing proteins arrive, is largely determined by the bulk concentration (Cp). Protein mixtures, such as serum or plasma, and moderately hydrophobic surfaces are the common systems that demonstrate the Vroman effect. In these cases, high molecular weight species, such as immunoglobulins and kininogen start to replace albumin and/or fibrinogen (Lee and Lee, 1998). The surface and the solution conditions determine the relative adhesion strengths, so determine whether the Vroman effect will be seen. Albumin binds strongly to plain polystyrene through dispersive forces on the partially unfolded protein. In general, it cannot be displaced by larger proteins such as vitronectin, fibronectin, or immunoglobulins, as the desorption rate of its domains is very small compared to the time over which enzyme-linked immunosorbent assays (ELISA) are performed (i.e. Kads-eff is near infinite). Also the binding energy for a domain of a larger protein is unlikely to be larger than the binding energy for the partially unfolded albumin. Thus its success as a blocking agent for ELISAs. Plasma modified polystyrene, on the other hand, has a much reduced binding energy for albumin leading to more rapid desorption rates and the opportunity for larger proteins to displace it.

28.4.2 Textured surfaces Roughness of a surface adds its own extra dimension to the problem. If the rough surface can be treated as chemically homogeneous and the roughness dimension is larger than the protein dimension, then roughness simply adds available surface area for adsorbing protein. This can be simply treated by allowing an increased capacity for adsorbed protein. However, roughness often is associated with changed surface chemistry, especially at molecularly abrupt changes in direction of the surface. Surface modification with ultrathin layers of polymers (often by plasma deposition) may alter the surface chemistry quite considerably. This alteration is seen as a change in surface free energy, especially the polar component, and as a change in protein adsorption behaviour. Thin layers of polymers may be mobile enough to show the types of behaviour associated with mobile surfaces. Recently there has been much effort devoted to the development and manufacture of `nanotextured' or structured surfaces (1 to 1000 nm features) (Boeckl et al., 1998; De Cupere et al., 2003; Li et al., 2004; Morimoto et al.,

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2004). Most proteins of interest (blood plasma proteins) range in dimension from a few nanometers in diameter for small globular proteins such as lysozyme (3 nm  3 nm  4.5 nm; Rovira-Bru et al., 2001) to over 15 nm for large globular proteins (>1,000,000 MW). Rod shaped proteins, such as fibrinogen (MW340K) may be up to 50 nm long and 5 to 9 nm in diameter (Hall and Slayter, 1959; Jung et al., 2003) and collagen monomer is approximately 300 nm long and 1.5 nm in diameter. Changes in surface chemistry, or roughnesses, of the same dimension as a protein will affect its adsorption characteristics. Molecular imprinting is a technique that allows molecular shape, with some specificity, to be made in the surface of a polymer. Shi and Ratner (2000) have imprinted a number of proteins (ribonuclease A, -lactalbumin, immunoglobulin G, bovine serum albumin, fibrinogen and HEW lysozyme) in a polysaccharide polymer (trehalose) backed by a plasma-coated film from hexafluoropropylene. As the trehalose was coated from aqueous solution onto the adsorbed proteins, on a mica base, it is likely that there was some hydrogen bond specificity as well as simple shape specificity. Indeed, they found specificity for lysozyme and RNaseA, but much weaker specificity for lactalbumin. The technique, though, has promise, especially if imprinting matrices, which will allow stronger hydrogen bonding or electrostatic attraction along with the dispersive forces, can be found. Ion implantation, nano-lithography, self-assembled monolayers and nanoparticle deposition have been used to generate nanometer and micrometer-sized textures on a variety of surfaces. Relatively large, regular surface features can be made by lithographic techniques. Some of these are aimed at modifying protein adsorption by changing the hydrophobicity/hydrophilicity of the surface. For example, by appropriate choice of feature size and shape, a relatively hydrophilic material, such as silica (Chibowski and Perea-Carpio, 2001) can be made to be `super' hydrophobic ± that is, water drops form a contact angle close to 180ë to the surface. Individual features tend to be large compared to proteins, so that each feature may behave towards the protein similarly to the normal surface. The gaps between features will behave as fluid-gas interfaces.

28.5 Non-specific adsorption to `soft' surfaces Some surfaces, especially those with adhered or bonded, long hydrophilic chains, allow some or partial molecular penetration whilst excluding large molecules. In these surfaces the outer chains are permanently hydrated, although the polymer chains may contain hydrophobic regions. The effective surface free energy is equal to or greater than that of water (73 mJ/m2), so that it is impossible to estimate free energy by the contact angle method. Water is bound to the polar/hydrogen bonding groups. There are possibilities for charge-charge, hydrogen bonding and dipole interactions within the polymer chains as well as with proteins and other biomolecules.

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A common example is the long, polyethylene glycol (PEG, PEO) brushes used to reduce total non-specific protein adsorption, a process also referred to as anti-fouling. The reduction in non-specific protein binding for certain PEG surfaces has been related to water structuring around the PEG terminal regions (Mrksich and Whitesides, 1996; Harder et al., 1998; Ostuni et al., 2001), such that it is energetically unfavourable for proteins generally to bind. The theoretical treatments are borne out, in general, by adsorption studies (Meng et al., 2004; Kenausis et al., 2000). Surfaces covered by dense films of short chain PEG are resistant to non-specific protein binding. Harder et al. (1998) have shown that the PEG-water structure is crucial to the effect and that subtle changes in the length of the PEG chain or the surface density of the PEG will change the overall structure and present a surface that will be more susceptible to protein binding. A reduced density in PEG chains on the surface may also expose the underlying material to protein adsorption. Dry methods of estimating uniformity of coverage (e.g. ESCA, XRD, SI-TOF-SIMS) may not detect holes that may occur when the film is hydrated. There currently does not exist a suitable method for detecting holes in films. Other polymers that have some potential include acrylamide, acrylic acid, charged polymers and polysaccharides. The effectiveness in prevention of nonspecific binding or, in the case of microfiltration, fouling is variable depending on the polymer used as the surface coat. The principle in common would appear to be the generation of an hydrophilic surface layer to which the major plasma proteins do not bind. Applications have been mainly for microfiltration membranes, but are becoming more common for biosensors and other applications where small molecules are required to diffuse through, as well as blood contacting materials and contact lenses. As all the current partially penetrable surfaces are hydrogels, they have effects on water structure at the surface. Lipid adsorption on these hydrogels may affect further protein adsorption, as well as the apparent surface free energy of the biomaterials.

28.6 Non-specific adsorption to penetrable surfaces and interfaces Hydrogels are more frequently being used as biomaterials or biomaterial surfaces. Most hydrogels contain more water than polymer (>50% water) and the average interchain distances are several nanometers. This means that penetration of macromolecules into the hydrogels is to be expected. Surface, or near surface adsorption of protein therefore, is only part of the story with regard to these materials. For many hydrogels (water less than 90% of gel), the effective pore size may be around that of the small to medium size plasma proteins, that is, the pores are of the order of 3 to 5 nm. The pore size will depend heavily on the cross-link density of the gel as well as the chemical composition of the polymer.

Protein adsorption to surfaces and interfaces

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Cell encapsulation has become a technique for protection of immunogenic cells from the host. -islets of Langerhans cells are the most commonly attempted cell, due to the potential for reversing type I diabetes. The encapsulation must allow adequate nutrient flow, especially glucose, whilst preventing the antibodies that destroyed the host's -islet cells from destroying the implanted cells. This gives an effective cut-off of below 180 kDa. Poly-NIPAAm (N-isopropylacrylamide) is a very interesting polymer for its thermal characteristics. As a film over a solid substrate, at 37 ëC it presents a moderately hydrophobic `surface' to the overlying solution. This enables it to bind cellular adhesion factors and for cells to grow on it. If the temperature is lowered to, say 20 ëC, a phase change occurs that causes the surface facing the solution to become relatively hydrophilic. This reduces the adhesion strength of the adsorbed protein layer and the layer of cells can be stripped off. Not all protein is removed with the cells. At the next level of penetrability are surfaces that allow large macromolcules to penetrate, but keep out very large molecules, cells, viruses and molecular assemblies. Applications for surfaces with these capabilities include encapsulation, implantable optical lenticles, artificial liver and biosensors for large molecules. Very deep interface layers have the potential to be used as extra-cellular matrix (ECM) biomimics. Layer-by-layer surfaces are made from alternating layers of polycationic and polyanionic polymers (Li et al., 2004; Muller et al., 2004; Yang et al., 2001; Jiang and Hammond, 2000; Caruso and Schuler, 2000). Charge-charge interactions stabilise these layer-by-layer constructs, but they are still hydrogels and display partial penetrability by proteins, as well as by small molecules, depending on the polymer chain chemistry, molecular weight, hydration and crosslinking (Muller et al., 2004; Ngankam et al., 2004; Sukhorukov et al., 2004; Yu and Caruso, 2003; Emoto et al., 2000).

28.7 Future trends As we become better able to model protein-surface interactions and to predict the structures of proteins under different conditions, we will be better able to understand the various factors involved in the complex process of biomaterials interacting with the host. The ultimate aim is to generate sufficient information that we can confidently predict the host's cell and tissue behaviour and be able to construct surfaces that control the host's responses at all levels. In order to achieve this, we will have to improve on our ability to construct surfaces at the molecular level. This should allow us to produce surfaces that bind proteins specifically and in specific conformations and orientations. The choice of surface will then be able to be made on the basis of desired performance: bonding to tissue (e.g. bone, muscle); overgrowth by a specific cell type (e.g. vascular endothelium); lack of cellular overgrowth, collagen

776

Surfaces and interfaces for biomaterials

deposition or non-specific protein deposition (e.g. biosensors); and finally, control of cellular growth and differentiation both spatially and temporally (tissue engineering). The tools for this advance, as far as those involved in protein adsorption are concerned, are: · Predictive models for protein interactions with surface molecules (Latour, 1999; Raffaini and Ganazzoli, 2004) in aqueous solutions containing multivalent salts and other small molecules. This will require increases in massively parallel computing techniques at the molecular modelling level, and improved approximation techniques for determining the protein interaction energies for all allowable conformations of protein and surface. · Models for dealing with complex biological solutions, mixtures of proteins, lipids salts, saccharides and other biochemical species. · Models for dealing with protein turnover on the surface. · Choice of surface type for application will be essential. These may vary from the hard surfaces all the way through to cell-like surfaces that allow molecular penetration, even by large proteins. · Advanced biomimicry will be essential. Surfaces must not appear to be foreign to the host. · Models for dealing with direct cell responses to surfaces, including mechanical and electrical effects, as well as the chemical effect. · There will be erodable surfaces for short-term applications or for tissue engineering applications that are not completely bio-mimicing, but allow for control of the healing and tissue regeneration response. The near future will be devoted to developing these tools and techniques. Longer term will require the integration of knowledge obtained through genomics, proteomics, cellomics and the physiome, about how the various cues control the host response and how cues will have to change with time to achieve the desired effects ± normally integration and a return to a functional tissue or organ.

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substratum surface free energy on growth and spreading of human fibroblasts in the presence and absence of serum proteins. Journal of Biomedical Materials Research 20: 773±784. Scotchford CA, Gilmore CP, Cooper E, Leggett GJ, and Downes S, 2002. Protein adsorption and human osteoblast-like cell attachment and growth on alkylthiol on gold self-assembled monolayers. Journal of Biomedical Materials Research 59: 84± 99. Scotchford CA, Ball M, Winkelmann M, Voros J, Csucs C, Brunette DM, Danuser G and Textor M, 2003. Chemically patterned, metal-oxide-based surfaces produced by photolithographic techniques for studying protein- and cell-interactions. II: Protein adsorption and early cell interactions. Biomaterials 24: 1147±1158. Shi H and Ratner B, 2000. Template recognition of protein-imprinted polymer surfaces. J. Biomedical Materials Research 49: 1±11. Slack SM and Horbett TA, 1992. Changes in fibrinogen adsorbed to segmented polyurethanes and hydroxyethylmethacrylate-ethylmethacrylate copolymers. Journal of Biomedical Materials Research 26(12): 1633±1649. Stenberg, E., et al., 1991. J. Coll. Interface Sci. 143: 513. Song D and Forciniti D, 2000. Effects of cosolvents and pH on protein adsorption on polystyrene latex: A dynamic light scattering study. Journal of Colloid and Interface Science, 22: 25±37. Sugawara T and Matsuda T, 1996. Synthesis of phenylazido-derivatized substances and photochemical surface modification to immobilize functional groups. Journal of Biomedical Materials Research 32(2): 157±164. Sukhorukov GB, Volodkin DV, Gunther AJ, Petrov AI, Shenoy DB and Mohwald H, 2004. Porous calcium carbonate microparticles as templates for encapsulation of bioactive compounds. J. Materials Chemistry 14: 2072±2081. Ta TC and McDermott MT, 2000. Mapping interfacial chemistry induced variations in protein adsorption with scanning force microscopy. Analytical Chemistry 72: 2627± 2634. Toth J, 2002. Uniform and thermodynamically consistent interpretation of adsorption isotherms. In Adsorption: theory, modeling and analysis, J Toth (ed.), Marcel Dekker, New York, pp. 1±103. van Oss CJ, 2002. Use of the combined Lifshitz-van der Waals and Lewis acid-base approaches in determining the apolar and polar contributions to surface and interfacial tensins and free energies. Journal of Adhesion Science and Technology 16: 669±677. Vogler EA, 1998. Structure and reactivity of water at biomaterial surfaces. Advances in Colloid and Interface Science 74: 69±117. Voros J, 2004. The density and refractive index of adsorbing protein layers. Biophysical Journal 87: 553±561. Vroman L and Adams AL, 1969. Findings with recording ellipsometer suggesting rapid exchange of specific plasma proteins at liquid/solid interfaces. Surface Science 16: 438 et seq. Wagner, MS and Castner, DG, 2004. Analysis of adsorbed proteins by static time-offlight secondary ion mass spectrometry. Applied Surface Science, 231±2: 366±376. Wassell DT and Embery G, 1996. Adsorption of bovine serum albumin on to titanium powder. Biomaterials 17: 859±864. Weber N, Bolikal D, Bourke SL and Kohn J, 2004. J Biomedical Materials Research

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Index

ablative laser techniques 48, 469, 728 absorption-based SPR 265 absorption coefficient 255 absorption spectra 256±7 AC load 234±5 acoustic load concept (ALC) 331±2 acoustic/piezoelectric sensors 324±5, 357±60 activated partial thromboplastin time (APTT) 96 active sites 15±17 active transport 90 addition polymerisation 29 adenosine diphosphate (ADP) 112, 113 adhesion 240±1, 242 work of adhesion 364 adhesion assays see cell adhesion assays adsorption 11±12 hydrogen on Pt surfaces 15±17 kinetics of 15±17, 767±8 non-QCM gravimetric methods 361±2 proteins see protein adsorption restructuring surfaces and 14±15 adsorption isotherms 766±7, 769 aggregation 257, 262±4 air 40 albumin 109, 418, 769±70, 772 inhibition of platelet aggregation 427±8 preblocking of carrier surface 123, 125 serum see serum albumin alginate hydrogels 657 alkanethiol based SAMs 44±5, 260±1 alpha granules 98 alternating polymers 30 alternative complement pathway 110, 112 amorphous biominerals 676±8 see also biomineralisation amorphous calcium carbonate (ACC) 673±4, 681±2, 683±5, 687±8

amorphous polymers 32±4 amphiphilic block copolymers (ABCs) 576±7 amplitude parameters 696±9, 701±2 aneurysms 581 angioplasty 577 `animal on a chip' technology 484 animal models 432 anionic surfactant (AOT) 515 anisotropic etching 143 anisotropy 235 anodisation 64 antiobiotic resistance 622, 624 antibiotics 625±7 anticoagulants 431±2 anti-fouling 774 anti-oncogenes 756 antithrombogenic materials 737±8 aperture control 119 APTMS-APC 312±13 aragonite 672±6 arginine-glycine-aspartate (RGD) tripeptide 451, 613, 645, 751 arithmetical mean deviation of a profile 697±8 arteries 638±41 nanoparticles for vascular tissue drug delivery 577±8 replacement arteries 642±3 artificial pancreas 88 aspirin 416 association rate constant 191±6 atactic polymers 30 atomic force microscopy (AFM) 12, 13, 50, 51±2, 89, 398±9, 709±11 cell adhesion assays using AFM tip 759±60 coupled with QCM 357 modes 710

Index nanoindentation 225±7, 228±9 protein adsorption 95 sharpness of tip 711 thin films 77±8 use of an AFM tip in DPN 69±70 atrazine 354 attenuated total reflectance Fourier transform infra-red (ATR-FTIR) spectroscopy 50 attenuated total reflectance infra-red (ATR-IR) spectroscopy 50, 51, 76±7 attenuated total reflectance (ATR) spectroscopy 54 Auger electron spectroscopy (AES) 11, 53±4 Auger electrons 204, 208±9, 210 automatic null ellipsometers 279±80 automatic photometric ellipsometer 280±2 automobile exhausts 22±3, 24 averaging parameters 697±8 avidin biotin complex 136±7, 259, 351 backflushing 526, 528 backpulsing 526, 528 backscattered (reflected) electrons (RE) 204, 205, 210 bacteriophage 346 band-pass filters 695±6 Berkovich pyramid 230, 231 Bessel functions 286 bFGF 657, 658 binding inhibition test scheme 186 bioactive glasses 548, 549±50 scaffolds 554±7 for lung tissue engineering 564±5, 566 in vitro effects on bone regeneration 557±8 in vivo performance of scaffolds 560 bioactive materials 548 interface reactions in solution 549±50 see also bioactive glasses biochips 118±19 biocompatibility 39±40, 93, 389±90, 574±5, 693±4, 722±3 biosensors 105±16 polyphenolic films 125±8 surface modification of polymers to enhance 719±44 bioconnective materials 390, 405±10 biodegradable synthetic materials see synthetic biodegradable materials biofilms 345±6, 619±36

783

clinical implant infection 625±7 consequences of formation on clinical implants 624±5 determinants of antibiotic resistance 624 epidemiology and costs of infections 620±1 formation by clinical implants 619 further research 630 membrane separation systems 511±13 microbiology of clinical implant infections 621±2 molecular mechanisms underlying formation 622±4 prevention of formation 627±30 biofouling 105, 390 design for non-biofouling surface 399±405 in membrane separation systems 508±13 fouling control 513±29 Bioglass 548, 550, 554, 560 bio-hybrid devices 100 biointerface 573±5 see also intravascular drug delivery systems and devices biologically-induced mineralisation 667 biomaterials 389±413 bioreactions on surfaces 389±91 classification of surfaces and interfaces 764±5 defining 693±4 need for 545±7 polymer surface properties and biomaterials applications 56±7 surface analyses 393±9 contact angle method 394±5 SPM 397±9 XPS 396±7 surface design 391±3 surface properties 39±40, 389±93 biomimicry 117 biomimetic modifications of membranes 520±2 biomineralisation 666±92 biologically-induced and organic matrix-mediated 667 control over crystal orientation 674±6 control over crystal structure 671±4 control over mechanical properties 686±7 control over morphology 676±86 amorphous biominerals 676±8 crystalline biominerals 678±86 organic macromolecules 667±71

784

Index

biomolecular interaction analysis 190±6 direct detection 191±4 indirect detection 194±6 bioresponse compromise 104 biosensors 103±49, 151±2, 183±4, 338 biocompatibility 105±16 blood interfacing 111±14 protein constituents 108±11 tissue interfacing 114±16 and bioresponse compromise 104 interfacial problems 105 limitations 104±5 materials interfacing strategy 116±19 membrane systems used in 119±37 microflows as surrogate, renewable barrier films 138±40 microfluidics and 140±6 QCM 339±47 and selectivity compromise 104 biotins 135±7, 259, 351 birefringence 283±4 high birefringence (HiBi) fibre 284±91 blending, polymer 31, 391±2, 516, 521 block copolymer micelle lithography 168, 732 block (segmented) polymers 30 blocking laws 502±3 blood 57, 414±46, 574 haemocompatibility of metals, ceramics and polymers 425±7 interfacing and biosensors 111±14 ISO 10993 requirements for testing medical devices 431±2 membranes and thin films contact with 92±3 biological events at the interface 93±9, 100 problems of medical devices in contact with 414±18 role of blood cells during flow 422±3 surface characteristics in relation to haemocompatibility and clinical applications 423±5 surface interactions 418±22, 738±9 surface treatment to improve haemocompatibility 427±31, 736±40 test models for blood flow devices 432±4 blood vessels natural 638±41 see also arteries; vascular tissue engineering Bode plots 134±5, 136

body centred cubic (BCC) structure 5, 7 bone 410, 411, 545±6 nanoindentation 236±8, 240 orthopedic implants 547±8, 586, 620, 630 regeneration 552±61 surgical procedures for repair 547±50 bovine serum albumin (BSA) 107, 313±14, 508±9 Bradley (DMT) model 226 Bragg's law 71±3 branching 31 bridging 134 bright-field-image mode 209 brush grafting 41, 454, 519, 774 Butterworth-van Dyke (BVD) circuit model 323±4, 329 cake filtration 502±3 calcite 672±4, 679±81, 682±5, 686±7 calcium carbonate 666, 669, 672±4, 682, 684 amorphous 673±4, 681±2, 683±5, 687±8 calcium phosphate 666 calibration 713 cancellous (trabecular) bone 237±8, 545 cantilevers 213±15 capillaries 115±16 carbon-coated heart valve 57 carbon dioxide lasers 726, 727 carbon monoxide oxidation in auto exhausts 22±3, 24 oxidation on Pt surface 24±5 carbon phase diagram 18±19 cardiopulmonary bypass (CPB) 422±3 caries, dental 409±10 carious lesions 239 carrier phase ultrafiltration (CPUF) 523 catalysis, heterogeneous 20±3 catechol 127 catheters 581 CVCs 620, 625, 626, 629 urethral 628±9 cell adhesion assays 745±62 cell de-adhesion vs cell adhesion 747 current methods 757±60 cell population methods 757±8 single cell methods 758±60 issues of interpretation 750±3 cell bonding to biomaterial surfaces 750±2 cell detachment during an assay 752±3

Index measurement objectives 746±50 cell population techniques 746±9 sensitivity of the assay 749±50 single cell techniques 749 mode of detachment 747±8 sources of variability 753±6 test material surface 753±4 variability in cells used as probes 755±6 cell aggregates 481±2 cell culture analogue (CCA) devices 484 cell deposition, controlled see controlled cell deposition cell encapsulation 100, 775 cell population techniques 746±9, 757±8 cell tethering 748 cells 466 adhesion to polymer surfaces 467 cell tissue ± biomaterial interaction 389±413 connecting tissues with biomaterials 405±10 design for non±biolfouling surface 399±405 surface analyses of biomedical materials 393±9 surface properties of biomedical materials 389±93 cellular activation at membrane and thin film blood interface 97±9, 100 guidance through surface cues see surface cues in vivo and in vitro cell interactions 466±8 plasma membrane 83, 84 protein adsorption and cell attachment 315±16 response to biomaterials 603 see also cell adhesion assays cellulose-based membranes 85 central vascular catheters (CVCs) 620, 625, 626, 629 centrifugation-based adhesion assay 757±8 ceramic scaffolds 553±4 ceramics 425±7 chain scission 44 Chandler loop blood circulation model 433±4 charged groups 515 Charnley replacement joints 547±8 chemical cleaning of membranes 528±9 chemical `gating' 119 chemical patterning 448, 449±51

785

chemical sensors, QCMs as 333±8 chemical surface modification 99, 151, 515 chemical synapses 459 chemical vapour deposition (CVD) 17, 19, 64 chemistry investigating chemistry of polymer surfaces 53±5 and topography 150±1, 200±2 chip-based controlled cell deposition 465, 468±79, 483±4 chitin 674±5 chitosan 47 Clark pO2 polarographic electrode 110 classical complement pathway 110, 112 cleaning membranes 528±9 surfaces 5 cleavage 5, 6, 7 clinical implants 619±36 consequences of biofilm formation 624±5 epidemiology and costs of infection 620±1 facilitation of microbial colonisation and biofilm formation 619 infection 625±7 diagnosis 625±6 treatment 626±7 microbiology of infections 621±2 molecular mechanisms underlying biofilm formation 622±4 prevention of biofilm formation 627±30 problems of contact with blood 414±17 shape and size 608±9 see also synthetic biodegradable materials; and under individual types of implant coagulation 113±14, 738±9 activation of coagulation system 95±6, 418±19 coagulation cascade 110, 111 coascervation 654 coated QCMs 335±8 coating 49 antifouling modification of membrane surfaces 515±16, 521±2 ECM molecules on surfaces 646±51 nanoindentation 235±6 surface modification of polymers to enhance biocompatibility 735±6 synthetic scaffolds 643, 644

786

Index

coccoliths 685±6 cohesion, work of 364 collagen 638±9 gels 651±2, 653 microfibrils 648, 650 colloidal-based fabrication techniques 168±9 colloidal metals 248±9 colorimetric detection 263, 264 colostrum 26 compact (cortical) bone 237±8, 545 competitive test scheme 186 complement system 96±7, 110±11, 112, 419±20 complete blocking 502 complex amplitude reflection coefficients 277 complex refractive index 275±6, 278 complex shear modulus 331 concentration polarisation 493, 497±8, 505 condensation polymerisation 29 conducting electropolymerised films 119±20, 130±7 confocal microscopy 218±19 conformational change 770±1 constant distance mode 211 constant interaction mode 211 contact angle 56, 311±12, 765 dynamic contact angle (DCA) analysis 322, 362±5 method for surface analysis 394±5 static 362 contact angle hysteresis 362±3 contact area 232, 233 contact-based instruments 694±5, 703±5, 713±14 contact depth 232, 233 contact printing 45±7, 164 see also photolithography contact stiffness 227, 229 continuum elastic models 230 control macromolecules 667±71 see also biomineralisation controlled cell deposition 465±92 future trends 483±4 in vivo and in vitro cell interactions 466±8 three-dimensional 465, 480±3 two-dimensional 465, 468±79, 483±4 coordination 22 copolymer micelle nanolithography 168, 732 copolymerisation 30

copper 14±15 corona discharge treatment 395 coronary stents 416±17 cortical (compact) bone 237±8, 545 Coscinodiscus wailesii 676±8 costs of infection 620±1 critical flux 507, 530 critical gel concentration (CGC) 36 critical sized bone defects 546 critical surface tension 364 cross-flow filtration systems 89 crosslinking 31, 44 crosslinkable hyaluronan hydrogels 656±7 crosslinked graft copolymers 41 crystal growth, controlling 17±20 crystal orientation, control over 674±6 crystal structure control over in biomineralisation 671±4 crystal surfaces and surface preparation 5±10 crystalline biominerals 678±86 multi-crystalline arrays 685±6 soluble additives 679±81 spatial constraint 681±5 see also biomineralisation crystalline polymers 30, 32±4 crystals, single 686±7 cube-corner geometry 230±1 culture senescence 755±6 cuprammonium process 85 Cuprophan 121 cured cements 407 cut-off wavelength 695 cyclic voltammetry 132±4 Dacron 643 dark-field-image mode 209 data analysis routines 229±35, 242 dead end filtration 88 Dean vortices 524±5 decellularised arteries 642 defects, structural 10 degree of polymerisation 30 dendrimers 577 dense granules 98 density functional theory 23±4 density of summits 702 dental enamel 238±9, 407±8 dental tissues 238±9, 240, 407±10, 411 dentin 238±9, 240, 408±10 deposition methods 17, 19, 61±4 deposition ratio 66 depth profiles 207

Index desalination 99±100 deuterium labelling 300±2 Dexon sutures 587 dextran 260 dialysis 493 haemodialysis 87±8, 98, 406 diamond 18±19 diamond-like carbon (DLC) coatings 117, 120, 215 diatoms 676±8 dielectric constant 250, 253±4, 255±7 diffraction image 209 diffraction techniques 11, 73±4 diffusion 90, 145±6 diode lasers 726, 727 dioxins 353±4 dip-and-dry technique 348 dip-pen nanolithography (DPN) 69±70, 157±8, 646 dipole moment 323 direct assay format 190±4 direct foaming 553±4, 556±7 direct imaging 203, 204 dissipation factor 328±9 dissociation rate constant 191±6 DMT (Bradley) model 226 DNA microcontact printing of 47 QCM biosensors 340±1 Drude model 250 drug delivery systems 89, 586 intravascular see intravascular drug delivery systems and devices drug-eluting stents 580±1, 735±6 durability 389±90 dynamic contact angle (DCA) analysis 322, 362±5 elastic modulus 227, 231±5 elastic recovery 233±4 elasticity 217±18 elastin 639, 641, 648 elastin-based materials for seeding 654, 655 electrical synapses 459 electrochemical deposition (electroplating) 64 electrochemical impedance spectroscopy (EIS) 134±7 electrochemical QCM (EQCM) 330±1, 356 electrodialysis 86, 99 electron-beam (e-beam) lithography 48±9, 152, 153, 156, 168

787

electron diffraction 73±4 electron microscopy 205±7, 208±11 electronic conduction 131, 134, 135 electronic noses (e-noses) 335 electronic speckle pattern interferometry (ESPI) 708 electro-optic self-nulling ellipsometer 280 electropolymerised films 119±20, 123±37 conducting 119±20, 130±7 non-conducting 123±30 electrostatic double-layer (EDL) force 647 electrostatic layer-by-layer deposition 68±9, 775 ELISA assays 97 ellipsometry 94, 189±90, 271±98, 356±7 fibre-based polarisation modulated 282±5 future trends 292 high birefringence fibre polarisation modulation ellipsometry 285±91 history of 278±82 instrumentation and optical elements 278 interaction of light with matter 274±8 measurement technique 288±91 polarisation of light and 271±2 principles and mathematical basis 272±4 embossing 45±6 hot embossing 144 Emiliania huxleyi 685±6 emphysema 546±7 enamel, dental 238±9, 407±8 encapsulation cell encapsulation 100, 775 drug delivery 87, 89 thin films 78±9 endothelial cell surface heparan sulphate (ESHS) 522 endothelial cells (EC) 639, 640, 641 endothelium 578±9 energetics 55±6 energy-filtering TEM 209 enthalpic energies 769 enthalpy 37±8 entropic energies 769 entropy 37±8 environment and hydrolytic degradation of PGA 595±6 for nanoindentation testing 241±2 environmental SEM (E±SEM) 210±11 enzymatic degradation 590±8

788

Index

enzyme histochemistry 602, 609±13 eosinophils 98±9 epithelial cells 466 erythrocytes 99, 112 ESCA (X-ray photoelectron spectroscopy) 11, 50, 51, 53, 396±7 etching 85, 143 laser etching 48, 469, 728 plasma etching 723, 724 evanescent field techniques 186±90 ex vivo tests 95 excimer lasers 726, 727±30 expanded polytetrafluoroethylene (ePTFE) 643 exponential data analysis 192 external cake fouling 501±2, 508±10 extracellular matrix (ECM) 466±7, 751 components 638±9 vascular cell-matrix interactions 640±1 in vascular tissue engineering 637±65 coating on surfaces 646±51 ECM-regulated delivery of therapeutic growth factors 657±8 seeding materials 651±7 extracorporeal circulation (ECC) 417±18 face centred cubic (FCC) structure 5, 6, 9, 14 faceting 5 Faraday's law 330 fastest decay autocorrelation length 702 feedback 287±8 ferritin 671, 672 fibre based polarisation modulated ellipsometry 282±5 high birefringence 285±91 polarisation maintaining behaviour 283±5 polarisation state control and polarisation modulation 283 fibre Bragg gratings 292 fibrillin 639 microfibrils 648, 650, 654, 655 fibrillin-1 640±1 fibrin 110, 111 gels 652 fibrinogen 315, 316, 343, 418, 433, 739 fibroblasts 728±9, 730 fibronectin 647±8, 770, 771 fibulin±5 639, 641 Fick's law 194 film phase thickness 277 film thickness 326±7 filters 695±6

flow chambers 757, 758 flow manipulation 523±7 flow test models 432±4 flow through cell for QCM 331 fluid conductivity measurement 145 fluorescence labelling 185, 186 fluorescence microscopy 218 foaming bioactive ceramics 553±4 bioactive glasses 554, 555±7 foaming agents 555±6 force spectroscopy 217±18 formic acid 15 fouling of membranes 499, 500±30 biofouling 508±13 control 513±29 factors affecting 503±8 mechanisms 500±3 Fourier transform infra-red (FTIR) microscopy 55 Fourier transform infra-red (FTIR) spectroscopy 54 free energy of crystal faces 17±18 free radicals 599±601 free surface energy 364±5, 765±6 free volume 34 Freon 555±6 Fresnel equations 252, 274±5 Fresnel reflection coefficients 274±5, 276 Freundlich isotherm 766±7 friction force microscopy 216±17 fringed micelle 33, 34 functionality 389±90 fused deposition modelling (FDM) 46 fusion of cell aggregates 481±2 gallium orthophosphate 329, 332 gamma irradiation 44, 591±2, 593 gas separation 86 gas sparging 526±7 gel-casting 554 gels gel layer in membrane separation 498±9 polymer 35±6 gene activation bioactive glasses and bone regeneration 557±8 and the lung 565±6 gene therapy 55±6 genetically modified organisms (GMOs) 340 Gibbs Free Energy 37±8 glass 289±90

Index bioactive glasses see bioactive glasses glassy polymers 32±4 global analysis 192 glycolide 586 gold 6±7, 326 colloidal 248±9 SAMs 258±9 graft coupling 733±5 graft polymerisation 31, 32 surface modification 40±3, 391, 452±3, 516±18, 733±5 grafts/transplants bone 547 organ transplantation 550±1 vascular 414±16, 581 vascular cells used to seed 646 graphite 6, 8, 18±19 grating coupler 188 growth factors, therapeutic see therapeutic growth factors Haber-Bosch process 20±1, 22, 26 haematocrit 422 haemocompatibility 423±31 metals, ceramics and polymers 425±7 surface modification to improve 427±31, 736±40 see also blood haemodialysis 87±8, 98, 406 HAP2 303±5 hard biological materials see mineralised tissues `hard surfaces' 764 protein adsorption 765±8 hardness 217±18, 230, 231±5 Hayflick limit 755 heart-lung machine 421 heart valves 416±17 heat conduction calorimeter (HCC) 357 helical screw thread flow promoters 525 heparin 49, 109±10, 417 biomimetic modification of membranes 522 coatings and haemocompatibility 428±9, 434±5 reducing effects of coagulation 95±6, 114 surface modification to enhance biocompatibility 737±8 hetero-bifunctional cross-linker 452±3 heterogeneity index (HI) (polydispersity) 32 heterogeneous assays 186 heterogeneous catalysis 20±3

789

hexa-tert-butyl-decacylene (C60H66) 14 hexagonal close packed (HCP) structure 5 high birefringence (HiBi) fibre 284±5 polarisation modulation ellipsometry 285±91 high coordination sites 22 high-energy treatments 43±4 high-pass filters 695±6 high-performance tangential flow filtration (HPTFF) 523 histological studies 602±9 homogeneous assays 185, 186 homogeneous membranes 121±2 homopolymerisation 29±30 hot embossing 144 Huber-Mises criterion 229±30 human serum albumin (HSA) 301±2, 314 humidity sensor 334 hyaluronan (HA) 641 crosslinkable HA hydrogels 656±7 HA supports 652±4 Hyaff-11 652±4 hybrid parameters 699±700, 702±3 hybridisation studies 185 hybridised dentin 408±10 hydrated polymer membrane 403±4, 405 hydrodynamic management 523±7 hydrogel based polymers 733 hydrogel membrane 520 hydrogels 35±6, 656±7, 774 alginate 657 crosslinkable hyaluronan 656±7 PEG-based 656 hydrogen adsorption 15±17 hydrophilicity 38, 57 antifouling modification of membrane surfaces 515±18 hydrophilic PDMS stamp 164 and protein adsorption 305±7, 314±15, 316±17 hydrophobicity 38, 766 hydrophobic effect of polymers 43 membranes and fouling 504 protein adsorption 307±12, 316±17, 769±70 hydroxyapatite (HA) 545 synthetic HA 548, 550 hydroxyl radicals 599±600 ICAM-1 (intercellular adhesion molecule-1) 578 ideal scaffold 552

790

Index

imaging 202±8 parameters and requirements of different methods 204±8 SPR 262, 264±5 immortal cell lines 755 immune system 96±7, 110±11, 112, 419±20 immunoglobulins 165, 167, 418 immunosensors 338 QCM 347±55 impedance approximation 331±2 impedance spectroscopy 134±7 implants see clinical implants in-situ microfabrication 146 in-stent restenosis 580, 735±6 in vitro circulation models 432, 433±4 in vivo monitoring 138±40 needle biosensors 138 open microflow for 138±40 indirect assay format 191, 194±6 inert surfaces 475 inflammatory cells 596±8, 599 inflammatory tissue reactions 114±15, 601±13 infra-red (IR) microscopy 54±5 infra-red (IR) spectroscopy 54, 75±7 inhibition tests 194±6 inhomogeneity, surface 768 ink-jet printing technique 46±7, 70±1 integrins 562, 640 interaction energies 769 interdigitated structures 132±4 interfaces, classification of 764±5 interferometer type waveguides 188 intermediate blocking 502 internal membrane fouling 501±2, 508±9 interpenetrating graft copolymers 40±1 interpenetrating polymer network (IPN) 392 intimal hyperplasia 415±16, 417 intravascular drug delivery systems and devices 573±84 biomaterials and biointerface 573±5 drug delivery systems 575 future trends 581±2 nanoparticles 575±9 stents 580±1 vascular grafts and catheters 581 intrinsic rejection coefficient 496 intrinsic transmission coefficient 496 inverse photoemission spectroscopy (IPES) 11 ion transport 92 ionic strength 505±6

irreversibly bound protein 767 ISO standards 722 ISO 10993 requirements for testing of medical devices 431±2 isotactic polymers 30 isotopic labelling 299, 300±2 jet impingement technique 757 JKR model 226 joint replacements 547±8, 620, 630 Jones calculus 274 Kanazawa-Gordon equation 327, 347 kinetics of adsorption 15±17, 767±8 Kretschmann configuration 251±2 kurtosis 699, 702 label-free screening 343 lactate 115 lactide 586 laminar flow patterning 478±9 laminin 454±5, 562 Langmuir adsorption isotherm 194, 766±7 Langmuir-Blodgett (LB) films 65±7, 69, 336 Large-T antigen construct 756 laser etching (laser ablation) (LAB) 48, 469, 728 laser profilometers 706, 714 lasers 725±30 lateral resolution 204±7 lateral transport control 122±3, 124, 125 latex piezoelectric immunoassay (LPEIA) 353 layer-by-layer (LBL) deposition 68±9, 775 Lennard Jones potential 214 leucocytes 97, 420±1 lift-off technique 143 limiting flux 498±9 limulus ameobocyte lysate (LAL) test 347 linear transformation 192 liquids, QCM measurements in 327±9 lithography 152±67, 171, 773 colloidal-based techniques 168±9 with photons, particles and scanning probes 152±8 soft lithographic techniques see soft lithographic techniques surface modification to enhance biocompatibility 730±2 see also photolithography

Index load-depth curves 227, 229, 231±3, 234, 240±1 long circulating particles 576±7 long-term implantation of biosensors 107 Love wave devices 358±9 low-density neuronal cultures 460 low-energy electron diffraction (LEED) 11 low molecular weight cut-off films 123±5 low-pass filters 695±6 lower critical solution temperature (LCST) 36, 37 lung 546±7 surgical procedures for repair 550±1 tissue engineering 561±7 luteinising hormone (LH) 135±7 lysosomal enzymes 591 lysozyme adsorption 299, 305±14 at C15OH/solution interface 309±12 at hydrophilic silicon oxide/solution interface 305±7 at hydrophobic OTS/solution interface 307±9 at PC monolayer/solution interface 312±14 Mac-1 integrin cell receptor 751 macrophages 97, 550, 591, 596, 599, 609 macroscopic regime 23±4 magnetic resonance imaging (MRI) 435 magnetite 666, 671±2 magnetosomes 671±2, 673 mask fabrication 470±1 mass transfer coefficient 498±9 mass transport 193 matrix devices 89 matrix effect 140 matrix templates 645±6 Maxon 588, 601, 610±11, 612±13 Maxwell-Garnett formula 255±6 mean free path 207 mean width of profile elements 699±700 mechanical properties, control over 686±7 melt-derived bioactive glasses 548, 550 scaffolds 554 membrane selectivity 497 membrane separation 493±542 applications 86, 499±500 biofouling 508±13 microbial fouling 511±13 proteins 508±10 concepts 495±9 factors affecting fouling 503±8

791

membrane materials and morphologies 503±4 operating conditions 507±8 solution environment 504±6 fouling control 513±29 antifouling modification of membrane surfaces 514±22 flow manipulation 523±7 membrane cleaning 528±9 optimisation of operating conditions 522±3 selection of membrane 513±14 fouling mechanisms and mathematical description of fouling 500±3 future trends 529±30 processes 86, 90, 99±100 membranes 83±102 applications 86, 87±9 for biosensor interfacing 116±19 conferred functional advantages 118±19 property requirements 118 biological events at the membrane blood interface 93±9, 100 activation of coagulation system 95±6 activation of immune system 96±7 cellular activation 97±9, 100 protein deposition or adsorption 93±5 blood material contact 92±3 characterisation 89±92 physical structure 89 transport processes 90±2 materials 83±7 membrane systems used in biosensors 119±37 conducting electropolymerised films 119±20, 130±7 non-conducting electropolymerised films 123±30 thick membrane films 119, 120±3 QCM biosensors 342±3 mesenchymal cells 466 mesenchymal stem cells (MSC) 640, 641 mesoscopic regime 23±4 metallic nanoparticles 255±7, 261, 262±4 metallic thin films 250±5, 257±62 metals, haemocompatibility of 425±7 methacryloyloxyethyl phosphorylcholine (MPC) polymers 49, 401±5, 406, 520±2

792

Index

micelles block copolymer micelle nanolithography 168, 732 fringed micelle 33, 34 micro total analytical systems (TAS) 142 microbial colonisation 619 microbiology of clinical implant infections 621±2 see also biofilms microbial films 107±8 microbial fouling 511±13 see also biofouling microcantilever techniques 759±60 microchannels 159±60, 161, 165±7, 476±8 microcontact printing (CP) 152, 159, 160±4 cell guidance through surface cues 456±8 controlled cell deposition 473±6 of DNA 47 microelectromechanical systems (MEMS) 108, 138, 140 micro-encapsulation 100, 775 microfabrication 730±2 in-situ 151 microfibrils 641, 648, 650 microfiltration (MF) 86, 116, 493±4, 501±3 microflows 138±40 microfluidic patterning 476±9 laminar flow patterning 478±9 using microchannels 159±60, 161, 165±7, 476±8 microfluidics 140±6 biosensor design and fabrication of microfluidic devices 142±4 detection 144±5 diffusion and sample preparation 145±6 in-situ microfabrication 146 microgravimetry 322±85 acoustic/piezoelectric sensors 324±5, 357±60 combination of QCM and other techniques 355±7 dynamic contact angle measurements 322, 362±5 non-QCM adsorption methods 361±2 QCM see quartz crystal microbalance thermal gravimetry 322, 360±1 microindentation 225 micro-injection moulding 144 microleakage 410

micromachining 47, 469 micrometre scale 200, 201 micro-moulding in capillaries (MIMIC) 159±60, 161, 165±7 micropipette methods 758±9 microporous material 184 microporous membranes 120±1 microscopic regime 23±4 microscopy 200±24 different concepts of imaging surfaces 202±4 electron microscopies 205±7, 208±11 future trends 219±20 imaging parameters and requirements of the different methods 204±8 nanoindentation 228±9 optical microscopies 206, 218±19 scanning probe microscopies 12, 13, 24, 77±8, 203±4, 211±18, 397±9, 709±13 see also under individual techniques microstamping on an activated polymer surface (MAPS) 163 micro-transfer moulding (TM) 159, 161 Miller indices 5 millimetre scale 201 mineralised tissues connecting with biomaterials 407±10, 411 nanoindentation 236±40 mode of detachment 747±8 modified bioactive foams 564±5, 566 modified cellulose membranes 85 molecular beam epitaxy (MBE) 17, 62 molecular electronics 60 see also thin films molecular imprinting 49, 185, 336±7, 340, 350, 773 molecular tilt elevation 74 molecular valve 519±20 molecular weight 591±2 nominal molecular weight cutoff 90, 495 mollusc nacre 674±6, 686, 688 Monocryl 589, 600±1, 605±6 monocytes 98 monofilament sutures 608 morphology control over in biomineralisation 676±86 membrane characterisation 89 membrane and fouling 504, 513±14 surface characterisation techniques 51±2

Index motor driven null ellipsometers 279±80 MTAC (methyltrialkyl cationic quaternary ammonium ion surfactant) 122 Muller matrix 274 multi-crystalline arrays 685±6 multifilament sutures 608 nacre, mollusc 674±6, 686, 688 nanobiotechnology 150±80 colloidal-based fabrication techniques 168±9 lithographic patterning 152±8 soft lithographic techniques 158±67 template-imprinted nanostructured surfaces 169, 170 nanofabrication 731 nanofiltration (NF) 86, 493±4 nanoindentation 225±47 data analysis 229±35, 242 hard biological materials 236±40 instrumentation 225±9 soft biological materials 240±1 thin films and coatings 235±6 nanometre scale 200, 201 nanoparticles 21 biological activity 25 as an intravascular drug delivery system 575±9 reservoir for systemic drug effects 576±7 vascular tissue delivery 577±9 metallic and SPR 255±7 applications 262±4 surface modification 261 nanotextured surfaces 772±3 natural processes 26 Navier-Stokes equation 142 Nd:YAG lasers 726, 727 near-field scanning optical microscopy 219, 709, 712±13 needle biosensors 138 neuronal networks, patterned 447, 458±61 neutron diffraction 73 neutron reflection (NR) 299±321 and deuterium labelling 300±2 effect of size of globular proteins on their adsorption 314±15, 316 lysozyme adsorption 305±14 peptide interfacial assembly 303±5 neutrophils activation 97 laser surface modification 728±9 platelet neurophil interactions 98

793

NIPAM (N-isopropyl acrylamide) 42 Nitinol 427 nitrogen-based fertiliser 20, 21 nitrogen fixation 26 nominal molecular weight cutoff (NMWC or MWCO) value 90, 495 non-biofouling surface 399±405 non-conducting electropolymerised films 123±30 non-contacting surface texture measurements 705±13, 713±14 non-selective mass measurement 333±8 non-specific adsorption 184±5 see also protein adsorption non-stimulative materials 390 nonthrombogenic materials 737±8 null ellipsometer 278±9 automatic 279±80 nulling technique 189±90 numerical aperture 205 nylon 6,6 membrane 146 observed rejection coefficient 496 observed transmission coefficient 496 octadecyl trichlorosilane (OTS) 307±9 Onsager coefficient 193 open flow planar membrane format 122±3, 124 open microflow 138±40 ophthalmologic tissues 606±8 optical microscopies 206, 218±19 optical techniques 12±14, 24±5, 94, 183±99 biomolecular interaction analysis 190±6 characterisation of surfaces 183±6 detection methods 186±90 surface texture measurement 705±8 optical tweezers 759 optical waveguide light spectroscopy (OWLS) 94, 356±7 organ damage monitoring 435 organ printing 481±3, 484 organ transplantation 550±1 organic macromolecules 667±71 organic matrix±mediated mineralisation 667 orientated crystal growth 674±6 orthopedic implants 547±8, 586, 620, 630 see also bone osmotic pressure 91±2 osteoporosis 546 oxidase-based biosensors 120±1 oxidation 40

794

Index

packing 34 packing density 34 paclitaxel 580 pancreas, artificial 88 parallel-plate flow chambers 757 parallel-resonant mode 323±4 Parylene 736 passivation of silicon surfaces 19±10 passive transport (diffusion) 90, 145±6 patterned neuronal cell culture 460±1 patterned neuronal networks 447 synaptic connections 458±61 patterning 45±9, 150±80 cell guidance through surface cues see surface cues colloidal-based fabrication techniques 168±9 lithographic with photons, particles and scanning probes 152±8 soft lithographic techniques 158±67 template-imprinted nanostructured surfaces 169, 170 vascular tissue engineering 646 peak and valley parameters 698±9 PEEK (polyetheretherketone) 143, 729±30 peeling 748 PEG-PLGA-PEG polymers 42 PEI(polyethyleneimine)PLGA block copolymers 45 penetrable surfaces and interfaces 774±5 pentadecyl trichlorosilane (C15OH) 309±12 peptide nucleic acids (PNA) 185 peptides bioactive and tissue engineering 562, 563 interfacial assembly 303±5 perfect terminations 7±8 permeability 90 permeable films 83±102 applications 87±9 biological events at film blood interface 93±9 blood material contact 92±3 materials 83±7 permeate flux 495±6 permeation parameters 89 pervaporation 86 pH 505±6 phase bias 285, 286±7 phase inversion 85 phase states of polymers 34±5 phase transitions 14, 34±5 phenolic monomers 128±9

phospholipids 334 phosphorylcholine (PC) 118, 737, 738 biomimetic modifications 520 coating and haemocompatibility 429±30, 739±40 neutron reflection at PC monolayer/ solution interface 312±14 phosphorylcholine zwitterionic layers 117 photochemical patterning 455±6 photoelectron spectroscopy (XPS) 11, 50, 51, 53, 396±7 photo-induced grafting 517±18 photolithography 45±7, 143 compared with soft lithography 159, 160 controlled cell deposition 468±70 patterning 152±3, 154±6, 168, 171, 731 patterning of chemical surface cues 454±5 photometric ellipsometer 280±2 physical cleaning 528 physical vapour deposition 61±3 piezoelectricity 322±3 acoustic/piezoelectric sensors 324±5, 357±60 see also quartz crystal microbalance (QCM) pile-up 228±9, 236 plane of slip (shear plane) 55 plasma-enhanced CVD (PECVD) 64 plasma etching 723, 724 plasma fractionation 88 plasma graft polymerisation technique 518, 733±5 plasma initiated polymerisation 723, 724 plasma lithography 155±6 plasma membrane 83, 84 plasma micropatterning 456 plasma separation (plasmapheresis) 88 plasma skimming layer 748 plasma surface modification techniques 43±4, 99, 100, 518, 723±5 plasmons 249 volume plasmons 248 see also surface plasmon resonance (SPR) platelets 98 activation 433 platelet neutrophil interactions 98 surface coagulation process 112±13 surface interactions of blood 421±2 platinum hydrogen adsorption onto 15±17

Index oxidation of carbon monoxide at surface 24±5 pluronic block copolymers (PBCs) 576±7 polarisation of light 271±2 control 278±82 see also ellipsometry polarisation modulated ellipsometry (PME) 281±2 fibre based 282±5 high birefringence fibre PME 285±91 polariser-compensator-sample-analyser (PCSA) ellipsometer 278±9 polarisability 255±6 polaronic conduction 131, 134, 135, 136 poling 64 polyacetylene 131 polybutyl methacrylate (PMB) 401±4 polycaprolactone (PCL) 46, 587 polycarbonate microporous membranes 120 polychlorinated dibenzo-p-dioxins (PCDDs) 354 polydimethylsiloxane (PDMS) 142±3, 144 moulds and stamps in soft lithographic techniques 159±60, 165, 472 stamps for microcontact printing 47, 159, 161, 162, 164, 457, 473 stencil patterning 479 poly-p-dioxanone (PDS) 587, 588±9, 610±11, 613 polyethacrylic acid (PEA) 43 polyethersulphone-PVP blend (PES-PVP) 516 polyethylene (PE) 30, 740 polyethylene glycol (PEG) 162 brushes 184, 774 hydrogels 154, 656 PEG modified nanoparticles 576 polyethylene oxide (PEO) 109±10, 114 polyglycolic acid (PGA) 587±8 in vivo biodegradation 603±5, 606±8, 613 role of enzymes in degradation 593±6 polylactic acid (PLA) 587, 613 polylactic-glycolic acid (PLGA) 644±5 foam scaffolds 613 polylactides 600 poly-L-lysine-g-polyethylene glycol (PLL-g-PEG) 163 poly-L-lysine-g-polyhistidine 42 polymer blending 31, 391±2, 516, 521 polymer brushes 41, 454, 519, 774 polymer gels 35±6

795

polymers 29±59 biocompatibility 722±3 coating of QCM electrodes 349±50 general properties of a biomaterial surface 39±40 haemocompatibility 425±7 MAPS 163 in medical applications 720±2 classification 720±1 compliance with USP and ISO standards 722 membranes 85±7, 116±18 see also membranes non-biofouling surface 399±405 polymer-solvent interactions 35±8 polymeric surface and surface-bulk difference 38 preparation of 29±32 solid state and structure 32±5 surface analysis 50±6 chemistry 53±5 energetics 55±6 morphology 51±2 surface modification see surface modification surface properties and biomaterials applications 56±7 polymethoxyethylacrylate (PMEA) 430±1 polymethylmethacrylate (PMMA) 142±3, 144, 311±12 poly-NIPAAm (N-isopropylacrylamide) 775 polyparaphenylene 131 polyphenolic films 125±8 biocompatibility and selectivity 125±8 polyphenol variants 128±9 with surfactant entrapment 129±30 polypropyl acrylic acid (PPA) 42±3 polypyrrole 131, 132±7 impedance spectroscopy at planar polypyrrole films 134±5 polyrosolic acid 128, 129 polysialylation 576 polythiophene 131 polyvinyl alcohol (PVA) 578±9 polyvinyl chloride (PVC) 121±2, 740 polyvinyl-pyrrolidone (PVP) 516 POPs stamp 164 pore size 86, 116±17 porosity 117 powder blasting 143±4, 469 pressure membrane separation and fouling 507 osmotic 91±2

796

Index

pressure driven membrane separation 493±4 pretreatment 505 primary pulmonary hypertension 546 prism couplers 188 profilometers, laser 706, 714 promoters 20 prosthetic joint implants 547±8, 620, 630 protein A 350±1 protein adsorption 151, 763±81 classification of biomaterials surfaces and interfaces 764±5 coating ECM molecules on surfaces 647±51 future trends 775±6 general rules of non±specific adsorption 768±73 dynamic binding 771±2 textured surfaces 772±3 `hard' surfaces 765±8 adsorption isotherms 766±7 kinetics of adsorption 767±8 surface inhomogeneity 768 at membrane and thin film interfaces 93±5 on metallic thin films 258±61 and neutron reflection 305±17 effect of size 314±15, 316 non-biofouling surface 399±405 penetrable surfaces and interfaces 774±5 and QCM 343±5 QCM compared with other techniques 356±7 `soft' surfaces 773±4 solid-fluid interface 467±8 protein films 241 protein patterning 151 proteins biocompatibility of biosensors 108±11, 112 deposition at membrane and thin film interface 93±4 fouling of membrane separation systems 508±10 interactions with surfaces and tissue engineering 561±3 orientation and conformation rearrangement 770±1 template imprinting 169, 170 temporal dependence of proteins deposited on human colostrum 26 proteoglycans 639 proton micromachining 47

pseudo-first-order reaction 192 Pseudomonas aeruginosa 623, 624 pulsatile flow 525±6 pulsed sample injection ultrafiltration 523 pyramidal indenter geometries 230 quality factor 326 quantification of surface texture 696 quartz crystal microbalance (QCM) 322±32 additional parameters 331±2 analytical applications 332±55 bioanalytical applications 338±55 biosensors 339±47 immunosensors 347±55 non-selective mass measurement 333±8 QCM as an analytical technique 332±3 combined with other techniques 355±7 contact angle measurement 365 electrochemical QCM 330±1, 356 measurements in liquids 327±9 sensors with other piezoelectric materials 332 QCM-D 322, 334, 344, 345, 356±7 radial flow chambers 757 radioactive isotope labelled proteins 94±5 Raman spectroscopy 11 random polymers 30 rapid prototyping 47 Rayleigh SAW sensors 357±8 reactive oxygen species 599±601 reagentless binding 135±7 rearrangement, surface 768 rebinding 193±4 reconstructed surfaces 8±10 red blood cells 99, 112 reflection by ambient-film-substrate system 276±8 by ambient-substrate boundary system 274±6 reflection absorption infra-red spectroscopy (RAIRS) 11, 76±7 reflection anisotropy microscopy (RAM) 12, 24±5 reflection anisotropy spectroscopy (RAS) 12±14 reflection high energy electron diffraction (RHEED) 74 reflectivity 706 neutron reflectivity 300±1

Index reflectivity profiles 252±4 reflectometric interference spectroscopy (RIfS) 189, 190 reflectometry 186±7, 189±90 refractometry 186±9 regenerative medicine 551±67 bone regeneration 552±61 ideal scaffold 552 tissue engineering of the lung 561±7 rejection coefficients 91, 496 relaxation processes 35 repeating units 30 replacement arteries 642±3 replacement joints 547±8, 620, 630 replica moulding (REM) 159, 161 residual stresses 233 Resolut membrane 588, 606 restenosis 577 in-stent restenosis 580, 735±6 restructuring 14±15 reverse osmosis (RO) 86, 116, 493±4 reversibility lysozyme adsorption 305±6, 307, 311, 315 surface spectroscopies 183±5 Reynolds number 141±2, 478 RGD (arginine-glycine-aspartate) tripeptide 451, 613, 645, 751 RNA III inhibiting peptide (RIP) 627 roller pump closed-loop blood circulation model 433±4 root-mean-square (RMS) of the deviation of the profile 698 root-mean-square (RMS) of the deviation of the surface 701 rotating analyser ellipsometer (RAE) 280±1 roughness 648, 649 cell adhesion assays 753±4 protein adsorption and 772±3 and thrombogenicity 424±5 rubber elasticity 34±5 rupture event scanning (REVS) 354±5 ruthenium 22±3, 24 salivary pellicle 241 Salmonella typhimurium 353 sample separation 145±6 sandwich immunoassay format 122±3 Sauerbrey equation 326, 327, 328, 330 scaffolds 480 rapid prototyping 47 regenerative medicine 551±67 bone regeneration 552±61

797

ideal scaffold 552 tissue engineering of the lung 561±7 vascular tissue engineering 642±5, 658 scanning electrochemical microscopy (SECM) 213 scanning electron microscopy (SEM) 52, 206, 207±8, 210±11, 228±9 scanning force microscopy (SFM) 206, 207, 208, 213±18 scanning near-field optical microscopy (SNOM) 219, 709, 712±13 scanning near-field photolithography (SNP) 154, 155 scanning probe contact printing 164 scanning probe lithography (SPL) 156±7 scanning probe microscopy (SPM) 12, 13, 24, 77±8, 203±4, 211±18, 397±9, 709±13 scanning tunnelling microscopy (STM) 12, 13, 24, 51, 77, 212±13, 398 imaging parameters and requirements of the technique 206, 207, 208 non-contacting method for surface texture measurement 709, 711±12 scanning tunnelling spectroscopy (STS) 213 sea urchins 678, 679, 680 larvae 681±2 second harmonic generation (SHG) 14 secondary electrons (SE) 204, 205, 210 secondary ion mass spectroscopy (SIMS) 50, 51 seeding biological seeding materials 651±7 vascular cells used to seed vascular grafts 646 selectivity polyphenolic films 125±8 surface spectroscopies 183±5 selectivity compromise, biosensors and 104 self-assembled monolayers (SAMs) 44±5, 67±8, 731 controlled cell deposition 474±6 patterning 152, 154, 155, 158±9, 161, 171 QCM immunosensors 351±3 surface functionalisation 258±9 semiconductor technology 17±20 senescence, cell 755±6 sensitivity adhesion assay 749±50 surface sensitivity of surface microscopies 207

798

Index

surface spectroscopies 183±5 series-resonant mode 323±4 serum albumin 770 bovine 107, 313±14, 508±9 human 301±2, 314 Sewell score system 602±5 shape of implants 608±9 shear horizontal-acoustic plate mode (SH-APM) devices 324±5, 357±8 shear horizontal polarised surface acoustic wave modes (SH-SAW) 357±9 shear plane (plane of slip) 55 shear rate 507±8 shear stress blood flow dynamics and surface interactions 432±3 defining close to the surface 748±9 signalling metabolites 623 silaffins 669 silica deposition vesicles (SDVs) 676±8 silanes 121, 154 silanisation 184 silica 666±7, 669±71 silicatein protein filaments 670±1 silicon 162, 289±90, 559 gene activation 559, 565±6 passivation of surfaces 19±20 reconstructed surfaces 9±10 silicon oxide 305±7 silicones, on microporous membranes 120±1 silk fibroin gel 674±5 silver 627 simple cubic (SC) structure 5 single cell techniques 749, 758±60 sink-in 228±9, 236 sintering 85 Sirolimus 580 skewness parameter 698±9, 701±2 slip, plane of (shear plane) 55 smooth muscle cells (SMC) 640 soft biological materials 240±1 soft lithographic techniques 158±67, 470±9 laminar flow patterning 478±9 microcontact printing see microcontact printing microfluidic patterning using microchannels 159±60, 161, 165±7, 476±8 stencil patterning 479 `soft' surfaces, protein adsorption to 773±4

sol-gel bioactive glasses 550 scaffolds 555±7 for lung tissue engineering 564±5 sol-gel films 290±1 sol-gel transition 36 soluble additives 668, 679±81 solutes, low molecular weight 108 solutions, feed 504±6 solvent-assisted micro-moulding (SAMIM) 159, 160, 161 solvent-polymer interactions 35±8 sparging, gas 526±7 spatial constraint 681±5 spatial parameters 699±700, 702±3 specific adsorption 763 spectroscopic ellipsometry 282 spectroscopy see under individual techniques spin-coating 63±4, 78 sponge-like polymer membranes 682, 684 spreading coefficient 364 sputtering 62±3, 207 stabilised Zeeman laser (SZL) 281 stability 183±5 stainless steel 426 standard blocking 502 standards 431±2, 722 Staphylococcus aureus 622, 623, 625, 626, 627, 628 Staphylococcus epidermis 622, 623, 625, 626, 627 stencil patterning 479 stents 433±4, 580±1, 582 drug-eluting 580±1, 735±6 problems of contact with blood 414±17 steric hindrance 519±20 stimuli-responsive grafts 41±3 Stokes parameters 273±4 streptavidin-biotin interaction 259, 351 stretching 85 structured unfolding 302, 307, 309±10 structure factors 73 stylus-based instruments 694±5, 703±5, 713±14 sum frequency generation (SFG) 14 superoxide ions 599±601 supported lipid bilayers (SLBs) 751±2 supramolecular structures 185 surface acoustic wave (SAW) devices 324±5, 357±60 surface analyses 393±9 surface area of implants 608±9 surface cues 447±64 chemical patterning 448, 449±51

Index patterning of chemical surface cues 454±8 surface functionalisation 451±4 synaptic connections in patterned neuronal networks 458±61 topographical patterning 448, 449, 450 surface degradation see synthetic biodegradable materials surface energy 36±7 surface enhanced Raman scattering (SERS) 54 surface free energy 364±5, 765±6 surface grafting 40±3, 391, 452±3, 516±18, 733±5 surface inhomogeneity 768 surface modification 40±9, 719±44 antifouling modification of membrane surfaces 514±22 biocompatibility 722±3 future trends 740±1 metallic nanoparticles 261 methods and their efficiency 391±3 planar, metallic thin films 257±61 polymers in medical applications 720±2 protein adsorption and 772±3 sol-gel bioactive glass scaffolds 564±5 synthetic scaffolds 643±5 techniques 723±40 coatings 49, 735±6 grafting 40±3, 391, 452±3, 516±18, 733±5 haemocompatibility 427±31, 736±40 high-energy treatments 43±4 lasers 725±30 lithography see lithography; photolithograpy; soft lithographic techniques oxidation by air 40 plasma 43±4, 99, 100, 518, 723±5 SAMs see self-assembled monolayers surface grafting 40±3, 391, 452±3, 516±18, 733±5 surface patterning see patterning surface modifying additives (SMA) 430 surface molecular imprinting 49, 185, 336±7, 340, 350, 773 surface patterning see patterning surface plasmon resonance (SPR) 94, 188±9, 248±70 applications 261±4 metallic nanoparticles 262±4 thin films 261±2

799

phenomenon 249±57 metallic nanoparticles 255±7 thin films 250±5 QCM and 355±6 surface functionalisation 257±61 metallic nanoparticles 261 thin films 257±61 surface pressure vs area isotherm 65±6 surface rearrangement 768 surface texture see topography surface X-ray diffraction (SXRD) 11 surfaces 3±28 active sites and kinetics 15±17 characteristics 14±15 classification 764±5 controlling crystal growth 17±20 defining 694 experimental approaches to real surfaces 24±5 experimental investigation 4±14 crystal surfaces and surface preparations 5±10 diffraction techniques 11 photoelectron and Auger electron spectroscopy 11 scanning probe and optical techniques 12±14 TPD 12 UHV 4±5 vibration spectroscopies 11 heterogeneous catalysis 20±3 insight into biological activity of 25±6 theoretical advances 23±4 surfactants 121±2 bioactive ceramic scaffolds 553±4 polyphenols with surfactant entrapment 129±30 surface charge modification 515 surgical procedures bone repair 547±50 lung repair 550±1 sutures 587, 589 in vitro degradation 591±601 tissue reaction 603±13 synaptic connections 458±61 syndiotactic polymers 30 synthetic biodegradable materials 585±618 chemistry of 587±9 in vitro degradation 589±601 role of enzymes 590±8 role of free radical and superoxide ions 599±601

800

Index

in vivo degradation and cell/biomaterial surface interaction 601±13 synthetic HA scaffolds 548, 550, 559 synthetic scaffolds 643 modified synthetic scaffold materials 643±5 tantalum 426 Taylor vortices 524 teeth 238±9, 240, 407±10, 411 temperature programmed desorption (TPD) 12 template imprinting 185 nanostructured surfaces 169, 170 templates, matrix 645±6 10% rule 235 terpolymerisation 30 telomerase gene (tert construct) 756 tethering, cell 748 texture aspect ratio 702 therapeutic growth factors ECM-regulated delivery of 657±8 TGF 1 639, 641, 657, 658 thermal evaporation 61±2, 78 thermal gravimetry (TG) 322, 360±1 thermokinetic analysis (TKA) 361 thermoplastic polymers 720±1 thermosetting polymers 720±1 thick membrane films 119, 120±3 thickness shear mode (TSM) devices 324±5 thin films 60±82 ellipsometry 273 reflection by ambient-film-substrate system 276±8 established deposition methods 61±4 future trends 78±9 molecular architectures 65±71 molecular organisation in 71±8 nanoindentation 235±6 permeable 83±102 applications 87±9 biological events at film blood interface 93±9 blood material contact 92±3 materials 83±7 SPR 250±5 applications 261±2 surface modification 257±61 three-dimensional controlled cell deposition techniques 465, 480±3 three-dimensional scaffolds 545±72 regenerative medicine 551±2 bone regeneration 552±61

tissue engineering of the lung 561±7 surgical procedures for bone repair 547±50 surgical procedures for lung repair 550±1 three-dimensional surface profile data 700±3 thrombin 95 thrombotic events 414, 415 thrombotic potential 95 thrombus growth 105, 113 tissue connecting with biomaterials 405±10 ECM 466 interaction between biomaterials and cell tissues 389±413 interfacing and biosensors 114±16 reactions to biodegradable materials 601±13 tissue engineering 100, 152, 466, 476, 484, 585, 613 of the lung 561±7 three-dimensional controlled cell deposition 465, 480±3 vascular see vascular tissue engineering tissue repair 114, 545±72 need for biomedical materials and implants 545±7 regenerative medicine 551±67 bone regeneration 552±61 tissue engineering of the lung 561±7 surgical procedures 547±51 bone repair 547±50 lung repair 550±1 titanium 426±7 topographical patterning 448, 449, 450 topography 693±716 biomaterials, surfaces and biocompatibility 693±4 and chemistry 150±1, 200±2 defining surface texture 694 filters 695±6 quantifying surface texture 696 surface measurement 694±5 techniques for surface texture measurement 703±13 contact measurements 694±5, 703±5, 713±14 non-contact measurements 705±14 three-dimensional surface data 700±3 traceability and calibration 713 two-dimensional surface data 696±700, 701 total complement 97

Index trabecular (cancellous) bone 237±8, 545 traceability 713 tracheal tubes 629 transfer ratio 66 transmembrane pressure pulsing 525±6 transmission coefficients 496 transmission electron microscopy (TEM) 206, 208, 209, 228±9 transplants see grafts/transplants transport control, membranes for 121±3, 124, 125 transport processes 90±2 Tresca criterion 229±30 tribology 202 tri-calcium phosphate ceramics ( -TCP) 548, 550 Trillium Bio-passive Surface (TBS) 428±9 Triton X-100 121, 122 tropoelastin 639, 641, 658 trypsin 593±5 tube inserts 525 tunica adventitia 638 tunica intima 638 tunica media 638 27MHz QCM 341±2 two-dimensional controlled cell deposition techniques 465, 468±79, 483±4 two-dimensional surface profile data 696±700, 701 two±layer model 308 ultrafiltration (UF) 86, 116, 493±4, 502 ultra-flat nanosphere lithography (UNSL) 169 ultra-high vacuum (UHV) 3±5, 17 uncoated QCMs 333±5 uniform layer model 302 upper critical solution temperature (UCST) 36 urethral catheters 628±9 USP standards 722 vacuum evaporation system 61±2 valley depths and peak heights 698±9 VAPG 656 vapour deposition chemical 17, 19, 64 physical 61±3 Varidase 596 vascular cells 639±40 interactions with matrix 640±1 used to seed vascular grafts 646

801

vascular development 641 vascular grafts 414±16, 581 vascular cells used to seed 646 vascular prostheses 416±17 vascular substitutes 642±3 vascular tissue delivery, nanoparticles for 577±9 vascular tissue engineering 637±65 biological seeding materials 651±7 bio-patterning 646 coating ECM molecules on surfaces 646±51 ECM-regulated delivery of therapeutic growth factors 657±8 future trends 658 matrix templates 645±6 modified synthetic scaffold materials 643±5 natural blood vessels 638±41 replacement arteries and scaffold materials 642±3 synthetic scaffolds 643 vascular cells used to seed vascular grafts 646 vaterite 672 VEGF-modified hydrogels 657 ventilator-associated pneumonia (VAP) 620, 629 vibration spectroscopies 11 vibrational energy levels 75 Vicryl 588, 592±3, 596±8, 599, 606 viscoelastic materials 235, 240, 242 QCM 331±2 Visking process 85 vitronectin 770 volatile organic compounds 335 volume plasmons 248 volumetric flux 91, 495±6 vortex mixing 524±5 Vroman effect 94, 468, 771±2 Washburn method 363 wave vector 250±2 wavenumbers 75 wet etching 468±9 wettability 423±4 wetting force 363 wetting tension 364 white blood cells 97, 112 see also under individual types white light interferometry (WLI) 706±7, 714 wound cleaning 596 wound closure 586

802

Index

see also sutures wound healing 115±16 WWW15 303±5 X-ray diffraction 71±3 X-ray photoelectron spectroscopy (XPS or ESCA) 11, 50, 51, 53, 396±7

Young-Dupre equation 364 Young equation 56, 765 Young interferometer 188 Young's modulus 218 YYY15 303±5 zeta potential 55±6