Surface Modification of Biomaterials: Methods, Analysis and Applications (Woodhead Publishing in Materials)

Surface modification of biomaterials

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Related titles: Cellular response to biomaterials (ISBN 978-1-84569-358-9) The response of cells to biomaterials is critical in medical devices. It has been realised that specific cell responses may be beneficial ± encouraging adhesion, healing or cell multiplication. Cellular response to biomaterials will examine the response of cells with a wide range of materials, targeted at specific medical applications. Chapters in the first section review cellular response to polymers and ceramics. A second group of chapters discuss cell responses and regenerative medicine for nerves, muscles and orthopaedic materials. The final set of chapters analyse the effect of surface chemistry and how it can be manipulated to provoke a useful cell response. Degradation rate of bioresorbable materials: prediction and evaluation (ISBN 978-1-84569-329-9) Bioresorbable materials could be employed to provide significant advances in drug delivery systems and medical implants. The rate of material degradation is critical to performance of both implants and the controlled release of drugs. Degradation rate of bioresorbable materials: prediction and evaluation addresses the practicalities of this subject in detail. The first section provides an overview of bioresorbable materials and the biological environment. Degradation mechanisms are reviewed in the second section, followed by bioresorption test methods in the third part. The fourth section discusses factors influencing bioresorption; finally clinical applications are reviewed. Surfaces and interfaces for biomaterials (ISBN 978-1-85573-930-7) This book presents our current level of understanding on the nature of a biomaterial surface, the adaptive response of the biomatrix to that surface, techniques used to modify biocompatibility, and state-of-the-art characterisation techniques to follow the interfacial events at that surface. Details of these and other Woodhead Publishing materials books can be obtained by: · visiting our web site at www.woodheadpublishing.com · contacting Customer Services (e-mail: [email protected]; fax: +44 (0) 1223 893694; tel.: +44 (0) 1223 891358 ext. 130; address: Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, UK) If you would like to receive information on forthcoming titles, please send your address details to: Francis Dodds (address, tel. and fax as above; email: francis.dodds@ woodheadpublishing.com). Please confirm which subject areas you are interested in.

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Surface modification of biomaterials Methods, analysis and applications

Edited by Rachel Williams

ß Woodhead Publishing Limited, 2011

Published by Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, UK www.woodheadpublishing.com

Woodhead Publishing, 525 South 4th Street #241, Philadelphia, PA 19147, USA Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi ± 110002, India www.woodheadpublishingindia.com First published 2011, Woodhead Publishing Limited ß Woodhead Publishing Limited, 2011 The authors have asserted their moral rights. 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 publisher cannot assume responsibility for the validity of all materials. Neither the authors nor the publisher, 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 Woodhead Publishing Limited. The consent of Woodhead Publishing Limited 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 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. ISBN 978-1-84569-640-5 (print) ISBN 978-0-85709-076-8 (online) The publisher's policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp which is processed using acidfree and elemental chlorine-free practices. Furthermore, the publisher ensures that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by Godiva Publishing Services Limited, Coventry, West Midlands, UK Printed by TJI Digital, Padstow, Cornwall, UK

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Contents

Contributor contact details

xi

Preface

xv

Part I Surface modification techniques 1

Surface modification of biomaterials by plasma polymerization

E. J. SZILI, R. D. SHORT and D. A. STEELE, University of South Australia, Australia and J. W. BRADLEY, University of Liverpool, UK

3

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9

Introduction An overview of plasma and plasma polymerization Plasma generation and system design Plasma parameters Intrinsic parameters Potential biomaterial applications Future trends in plasma polymers Sources of further information and advice References

3 4 4 8 13 17 26 29 29

2

Surface modification of biomaterials by covalent binding of poly(ethylene glycol) (PEG)

39

Introduction Principles and methods Technologies and applications Conclusions and future trends References

39 40 51 52 54

A. RHODES, S. S. SANDHU and S. J. ONIS, BioInteractions Ltd, UK 2.1 2.2 2.3 2.4 2.5

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Contents

3

Surface modification of biomaterials by heparinisation to improve blood compatibility

X. ZHAO and J. M. COURTNEY, University of Strathclyde, UK

3.1 3.2 3.3 3.4

56

Introduction Bioactive molecule: heparin Blood±biomaterial interaction Surface modification by heparinisation for improved blood compatibility Future trends in heparinisation of biomaterial surfaces References

62 73 74

Surface modification of biomaterials by peptide functionalisation

78

4.1 4.2 4.3 4.4 4.5 4.6 4.7 4.8

Introduction Peptides and peptide functionalisation of surfaces Defining the biomaterial surface Peptide functionalised surfaces Non-covalent peptide functionalisation by self-assembly Spatial control of peptide functionality Conclusions References

78 78 81 81 88 91 95 95

5

Metal surface oxidation and surface interactions

102

Surface oxides in metallic medical devices: the scenario Titanium oxides on Ti implants: from crystallographic structure to the theoretical study of the atomistic surface structure and behaviour Technologies for tailoring Ti oxides on titanium Future trends References Appendix A: Materials and methods for unpublished results Appendix B: Abbreviations and symbols

102

3.5 3.6

4

L. S. BIRCHALL, H. QU and R. V. ULIJN, University of Strathclyde, UK

5.1 5.2 5.3 5.4 5.5 5.6 5.7

6

L. DE NARDO, G. RAFFAINI, F. GANAZZOLI and R. CHIESA, Politecnico di Milano, Italy

105 121 135 137 141 142

Surface modification of biomaterials by calcium phosphate deposition

143

Introduction Basic methods and applications

143 144

J. A. JUHASZ and S. M. BEST, University of Cambridge, UK

6.1 6.2

56 58 59

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6.3 6.4 6.5 6.6

Strengths and weaknesses Future trends Sources of further information and advice References

156 157 158 159

7

Biomaterial surface topography to control cellular response: technologies, cell behaviour and biomedical applications

169

Introduction Defining micro and nano Manufacturing surface topography How surface topography affects cell behaviour Technologies and potential applications Tissue regeneration Current issues and future trends Acknowledgements References

169 170 170 175 182 186 190 192 192

V. R. KEARNS, University of Liverpool, UK and R. J. McMURRAY and M. J. DALBY, University of Glasgow, UK 7.1 7.2 7.3 7.4 7.5 7.6 7.7 7.8 7.9

Part II Analytical techniques and applications 8

Techniques for analysing biomaterial surface chemistry

J. YANG and M. R. ALEXANDER, The University of Nottingham, UK 8.1 8.2 8.3 8.4 8.5 8.6 8.7

Introduction X-ray photoelectron spectroscopy (XPS) Time of flight secondary ion mass spectrometry (ToF SIMS) Sample preparation and handling Sources of further information and advice Acknowledgements References

9

Techniques for analyzing biomaterial surface structure, morphology and topography

N. S. MURTHY, Rutgers ± The State University of New Jersey, USA

9.1 9.2 9.3

Introduction Surface morphology and topography Surface structure and spatial distribution

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205 207 219 227 228 229 229

232

232 233 245

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Contents

9.4 9.5 9.6

Energetics Future trends References

10

Modifying biomaterial surfaces to optimise interactions with blood

248 250 253

A. DE MEL, Y. RAFIEI and B. G. COUSINS, University College London, UK and A. M. SEIFALIAN, University College London, UK and Royal Free Hampstead NHS Trust Hospital, UK 10.1 10.2 10.3 10.4 10.5

Introduction Physicochemical modification Biofunctionalisation Conclusions References

11

Modifying biomaterial surfaces with bioactives to control infection

255 263 268 271 273

H. J. GRIESSER, K. VASILEV, H. YS and S. A. AL-BATAINEH, University of South Australia, Australia 11.1 11.2 11.3 11.4 11.5 11.6

Introduction Plasma-based strategies for combating device-related infections Plasma polymers with incorporated metal nanoparticles or ions Covalent immobilisation of antibacterial molecules Future trends References

12

Modifying biomaterial surfaces to optimise interactions with soft tissues

J. GOUGH, University of Manchester, UK 12.1 12.2 12.3 12.4 12.5 12.6 12.7 12.8 12.9

Introduction Surface modification Surface modification Surface modification Surface modification Surface modification Conclusions Future trends References

of of of of of

biomaterials biomaterials biomaterials biomaterials biomaterials

255

for for for for for

the liver the kidney tendons/ligaments skeletal muscle skin

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284 287 290 294 304 305

309 309 309 312 314 316 319 321 321 322

Contents

13

Modifying biomaterial surfaces for the repair and regeneration of nerve cells

M. A. MATEOS-TIMONEDA, and J. A. PLANELL, The Institute for Bioengineering of Catalonia (IBEC), Spain and E. ENGEL, Technical University of Catalonia, Spain

ix

325

13.1 13.2 13.3 13.4

Introduction: the nervous system Surface properties and their effect on neural cells Future trends References

14

Modifying biomaterial surfaces to control stem cell growth and differentiation

344

14.1 14.2 14.3 14.4 14.5 14.6 14.7

Introduction Surfaces for stem cell expansion Surfaces for directing stem cell differentiation Surfaces for maintaining the differentiated cell phenotype Future trends Acknowledgements References

344 346 350 356 359 360 360

15

Modifying biomaterial surfaces to optimise interactions with bone

K. H. SMITH, Plataforma de Nanotechnologia, Spain and J. W. HAYCOCK, University of Sheffield, UK

R. L. SAMMONS, University of Birmingham, UK

15.1 15.2 15.3 15.4 15.5 15.6 15.7

Introduction Joint replacements: successes and challenges Bone graft substitute materials: successes and challenges Percutaneous bone-anchored devices Summary and conclusions Sources of further information and advice References

Index

325 329 334 338

365 365 368 375 382 390 391 391 401

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Contributor contact details

Chapter 2

(* = main contact)

Editor Dr Rachel Williams Department of Eye and Vision Science Institute of Ageing and Chronic Disease University of Liverpool First Floor Duncan Building Daulby Street Liverpool L69 3GA UK E-mail: [email protected]

Chapter 1 Dr Endre Szili, Professor Rob Short and Dr David Steele* Mawson Institute University of South Australia GPO Box 2471 Adelaide, SA 5001 Australia E-mail: [email protected]

Dr A. Rhodes*, Dr S. S. Sandhu and Mr S. J. Onis BioInteractions Ltd Science and Technology Centre Earley Gate Whiteknights Road Reading RG6 6BZ UK E-mail: [email protected]

Chapter 3 Professor Xiaobin Zhao* and Professor James M. Courtney Bioengineering Unit University of Strathclyde Glasgow G4 0NW UK E-mail: [email protected]; [email protected]

Professor James Bradley Department of Electrical Engineering and Electronics University of Liverpool Liverpool L69 3GJ UK ß Woodhead Publishing Limited, 2011

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Contributor contact details

Chapter 4

Chapter 7

Louise S. Birchall,* Honglei Qu and Rein V. Ulijn Department of Pure and Applied Chemistry University of Strathclyde 295 Cathedral Street Glasgow G1 1XL UK E-mail: [email protected] [email protected] [email protected]

Victoria R. Kearns* Institute of Ageing and Chronic Disease University of Liverpool Duncan Building Daulby Street Liverpool L69 3GA UK E-mail: [email protected]

Chapter 5 Dr Luigi De Nardo*, Professor Fabio Ganazzoli, Dr Giuseppina Raffaini and Professor Roberto Chiesa Politecnico di Milano Department of Chemistry, Materials and Chemical Engineering `Giulio Natta' Via Mancinelli 7 20131 Milan Italy E-mail: [email protected] [email protected] [email protected] [email protected]

Chapter 6 Dr Judith A Juhasz* and Professor Serena M. Best Cambridge Centre for Medical Materials Department of Materials Science and Metallurgy University of Cambridge Pembroke Street Cambridge CB2 3QZ UK E-mail: [email protected]

Rebecca J. McMurray and Matthew J. Dalby The Centre for Cell Engineering Joseph Black Building University of Glasgow Glasgow G12 8QQ UK E-mail: [email protected]

Chapter 8 Jing Yang and Morgan R. Alexander* Laboratory of Biophysics and Surface Analysis School of Pharmacy The University of Nottingham Nottingham NG7 2RD UK E-mail: [email protected]

Chapter 9 N. Sanjeeva Murthy New Jersey Center for Biomaterials Rutgers ± The State University of New Jersey 145 Bevier Road Piscataway, NJ 08854 USA E-mail: [email protected]

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Contributor contact details

Chapter 10 Achala de Mel, Yasmin Rafiei, Dr Brian G. Cousins and Professor Alexander M. Seifalian* Centre for Nanotechnology and Regenerative Medicine Division of Surgery and Interventional Science University College London Pond Street London NW3 2QG E-mail: [email protected]

Chapter 11 Hans J. Griesser* Ian Wark Research Institute University of South Australia Mawson Lakes SA 5095 Australia E-mail: [email protected] Krasimir Vasilev Mawson Institute University of South Australia Mawson Lakes SA 5095 Australia Hardi Ys Ian Wark Research Institute University of South Australia Mawson Lakes SA 5095 Australia and University of Tadulako Palu Indonesia

xiii

Sameer A. Al-Bataineh Mawson Institute University of South Australia Mawson Lakes SA 5095 Australia

Chapter 12 Dr Julie Gough School of Materials University of Manchester Grosvenor Street Manchester M1 7HS UK E-mail: [email protected]

Chapter 13 Dr Miguel Angel Mateos-Timoneda* and Professor Josep Anton Planell CIBER en Biomateriales, Bioingenieria y Nanomedicina The Institute for Bioengineering of Catalonia (IBEC) Parc CientõÂfic de Barcelona Edifici HeÁlix Baldiri Reixach 15-21 Barcelona, 08028 Spain E-mail: [email protected] [email protected] Professor Elisabeth Engel Biomaterials, Biomechanics and Tissue Engineering Group Department of Materials Science and Metallurgy Technical University of Catalonia, Avda Diagonal 647 Barcelona, 08028 Spain E-mail: [email protected]

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Contributor contact details

Chapter 14

Chapter 15

Katherine H. Smith Plataforma de Nanotechnologia Parc CientõÂfic Barcelona Baldiri Reixac 10-12 08028 Barcelona Spain E-mail: [email protected]

Dr Rachel L. Sammons School of Dentistry University of Birmingham St Chad's Queensway Birmingham B4 6NN UK E-mail: [email protected]

John W. Haycock* Kroto Research Institute Department of Engineering Materials University of Sheffield Broad Lane Sheffield S3 7HQ UK E-mail: [email protected]

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Preface

Many materials that are used in medical devices are well-established materials used in other areas of manufacturing. They therefore have well-defined properties in terms of both bulk properties and surface characteristics. This has allowed the appropriate choice of materials for specific devices with appropriate functional mechanics. However, frequently these materials do not have the appropriate surface properties essential to control the interaction of the device with the biological environment and thus may lead to clinical failure due to adverse biological responses. This has led to major programmes in surface modification of biomaterials to enhance their biological interactions. Surface modification of materials is a broad subject with many different considerations. For example, it is necessary to consider the properties of the materials to be modified, the surface characteristics required, the stability of the modifications, the application of the end product and the practicalities of the process. In some instances, surface chemical modification is advocated, whereas in others surface morphological changes are suggested. It is important to consider, however, that the routes to provide surface chemical modifications may also cause surface morphological changes and vice versa. An obvious result from this point is that it is essential to characterise all surfaces after surface modification so that all changes are fully defined. Considering that the biological interactions will be occurring at the atomic and molecular level, it is important to characterise the surface properties right down to the atomic- and nano-scale. Some surface modification processes are designed to incorporate specific functional groups, for example amine, hydroxyl, or methyl groups, with the aim of influencing the non-specific protein interactions with the surface and thus influencing the cellular interaction. Other techniques are designed to incorporate biological or biomimetic molecules, for example peptides or heparin, with the aim of providing a surface active layer that will interact directly with the cells. For metal surfaces, in particular, there are techniques designed to increase their passivity, such as oxidation, or enhance their bioactivity by coating with ceramics. Careful consideration is needed to choose the correct surface

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Preface

modification process to produce the surface characteristics required for the particular application and dependent on the properties of the materials being modified. In this book experts in their fields have provided detailed chapters on a range of surface modification techniques including plasma polymerisation, covalent binding of polyethylene glycol (PEG), heparinisation, peptide functionalisation, oxidation, calcium phosphate deposition and surface topographical modification. In many of these chapters there is practical advice on how to achieve optimised surfaces and a review of the current knowledge on how these influence biological interactions. Furthermore, all chapters provide an insight into future trends in the areas. Two key chapters (Chapters 8 and 9) have been put together to provide excellent reviews and practical advice on the surface characterisation methods for surface chemistry, structure and morphology. These provide important information on a range of surface characterisation techniques and will help the reader to choose the appropriate technique to use to define the surface properties dependent on the bulk material properties and the modification performed. The biological response that is required will clearly be dependent on the specific application for which the material is being designed and there may well be more than one route to obtain the required surface characteristics for the specific application. Although it is not possible to cover all application areas, in the second part of this book a range of key applications are addressed by authors with expertise in each area. This includes interactions with blood, soft tissues, nerve cells, stem cells, bone and the control of infection. In each case the authors review the role of surface modification in their specific application area and the performance of those currently under investigation. Furthermore they discuss where they think future trends are likely to take this research. It is clear from the research presented in this book that some surface modifications are well established and accepted in clinical practice, whereas others are still in the early stages of research. A greater understanding of how specific surface characteristics can be used to optimise the biological interactions with implantable biomaterials has the potential to enhance their clinical performance and demonstrates the importance of this field of research. Rachel Williams

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Surface modification of biomaterials by plasma polymerization E . J . S Z I L I , R . D . S H O R T and D . A . S T E E L E , University of South Australia and J . W . B R A D L E Y , University of Liverpool, UK

Abstract: The use of low pressure, low power plasma established in an atmosphere of precursor monomer(s) affords the deposition of functionalized films. Such films can retain the functional group and molecular structure of the starting compound(s) in an ultrathin `plasma polymer' deposit. A basic introduction to plasma and the modulation of some of the important physical and chemical properties will identify some of the key relationships between the plasma physics and the chemistry of the resultant plasma polymer films. There then follows an examination of where plasma polymerized films have found application in the area of biomaterial preparation and, in summary, identification of the likely direction of future developments. Key words: biomaterials, biomolecules, cell adhesion, plasma polymerization.

1.1

Introduction

The use of low pressure and, most recently atmospheric, low power plasma established in an atmosphere of precursor monomer(s) affords the deposition of functionalized films. Such films can retain the functional group and molecular structure of the starting compound(s) in an ultrathin `plasma polymer' deposit. Most commonly, surfaces are functionalised with oxygen (O), nitrogen (N), and O and N functional groups such as carboxylic acids, alcohols, ethers, amines and amides. This chapter outlines this surface engineering technology and highlights the importance that it has developed. In the first instance a basic introduction to plasma, how it is generated, and the modulation of some of the important physical and chemical properties is provided. In doing so, some of the relationships between the physics of low pressure, low power (radio-frequency) plasma discharges of O and N functionalized monomer(s) and the chemistry of the resultant plasma polymer films will be explored. There then follows, using selected examples, an examination of where plasma polymerized films have found, or have the potential to find biomedical application. We have structured this section according to the physiochemical characteristics of the plasma

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Surface modification of biomaterials

polymer that is being exploited; although rarely is control over one character manipulated to the exclusion of all others. In summary, we identify the likely direction of future developments.

1.2

An overview of plasma and plasma polymerization

Recent years have seen an increased interest in the preparation of functional, organic ultrathin films deposited by plasma. This technology is seen as the method of choice for modifying the surfaces of materials for biomedical and tissue engineering applications where the bulk mechanical properties have dictated the choice of material; however, the surfaces of such materials continue to elicit an undesirable biological-biomaterial response (Griesser et al., 1994; Chan et al., 1996; Daw et al., 1998; Barry et al., 2005; Siow et al., 2006; Zelzer et al., 2008). This is especially true as emerging technologies, which take advantage of events that occur at the nanometre and/or micrometre scale-length have developed (Morgenthaler et al., 2008; Robinson et al., 2008; Walker et al., 2009). For example, in devices utilizing microfluidic channels, the ratio of surface to bulk material is extremely high, hence the property of the surface in relation to the efficacy of the technology is paramount. Thus, in microfluidics, surface functionalization of microchannels can be utilized to both inhibit unwanted surface events and promote others (Hiratsuka et al., 2004; Sibarani et al., 2007). One particular feature of the plasma deposition of organic films, often referred to as `plasma polymerization', that makes these technologies particularly attractive is that such films can be deposited onto a wide range of different substrates without the need for special surface preparation prior to deposition. This provides a substantial advantage over other techniques where specific substrate chemistry is required, e.g. gold for thiols or oxidized surfaces for silanes; however, substantial commercial development of this technique has not yet followed: poor understanding of plasma systems, their operation, the chemical processes responsible for plasma polymer film growth coupled with the blanket description of these surfaces as biocompatible, based mainly upon short term in vitro studies, have restricted the field.

1.3

Plasma generation and system design

Historically the scientific investigation of plasmas began in the late 19th century when Sir William Crookes described a DC discharge in an evacuated column as `the fourth state of matter'; yet, it was Irving Langmuir in 1929, who first defined an ionized gas using the term `plasma'. However, the first industrial application of plasma processing occurred with the development of the modern integrated circuit first developed in the 1950s. Today anisotropic plasma etching allows patterning of integrated circuits (ICs) and the semiconductor industry,

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Surface modification of biomaterials by plasma polymerization

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with its desire for ever smaller features on microchips, has proven to be one of the main driving forces behind the research into plasma processing (Manos and Flamm, 1989; Flamm, 1996; Chang and Coburn, 2003). This research has, in turn, led to the almost ubiquitous implementation of plasma and plasma processing technology throughout the manufacturing sector. In addition to IC manufacture plasma coatings find application for modified adhesion ± both increasing and decreasing this property, scratch, corrosion and wear resistance (Wohlrab and Hofer, 1995; Villermaux et al., 1996; Benitez et al., 2000; Forch et al., 2007). Low temperature plasma surface modification is particularly attractive when applied to polymeric materials where the opportunity to modify surface properties such as wettability, adhesion, permeability and biocompatibility without thermal damage exists (Liston et al., 1993; Chan et al., 1996). Generation of low temperature plasma is achieved by use of DC or AC fields at reduced pressures, typically in the region of 1 to 100s of mTorr (0.133 to 10 s Pascals), such that ionization can occur at reasonable power inputs. Methods of excitation include DC, radio frequency (RF at 13.56 MHz) and microwave (MW at 2.45 GHz) and this power source may be coupled directly, with the electrodes within the plasma chamber, indirectly with the electrodes external to the chamber or by combination (Roth, 1995). For DC plasma the coupling is typically conductive between two electrodes and, depending on the application, a range of conducting materials may be used for electrode fabrication. With RF sources, the power, coupled to the electron current, can be capacitive or inductive. RF power coupling offers some significant advantages over DC and AC sources for industrial applications: · RF plasma can process insulating materials without sputtering of the electrodes and so, can be used for deposition from organic monomers. · Since the RF power is deposited in the plasma by displacement rather than particle currents, it is easier to couple through the chamber wall resulting in less ion and electron bombardment of electrodes. · In general, RF-generated plasma are more stable with electrons with higher temperatures for the same densities than an equivalent DC or AC plasma. This can be beneficial where an increased number of free radicals, plasmachemical reactions or dissociation and ionization reactions are desired. When considering the preparation of plasma polymer films, much of the previous research has been performed with RF power, typically 13.56 MHz, in glass reactor vessels with external coil-configurations, or bands to couple the power, although many studies with capacitively coupled, internal electrodes have also been undertaken (Griesser and Zientek, 1993; O'Toole et al., 1995; Beck et al., 1996; Dai et al., 1997; Alexander and Duc, 1998). Early designs were based on glass tubes which were either purpose made with ports for the vacuum, monomers and any diagnostic equipment, or utilized lengths of glass

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Surface modification of biomaterials

tube supplied `off the shelf' as high pressure column components. These allglass vessels are costly to design and produce and necessitated the use of an external driven coil. With glass tubes, readily available in a range of lengths and diameters, the various ports are incorporated into the flanges specifically constructed to enclose the vessel. Typically, these flanges are fabricated from conducting materials such as stainless steel or aluminium; however, insulating materials such as nylon can be used. Figure 1.1 illustrates the key components of such a plasma system ± the glass vessel is evacuated by a vacuum pump with an inline liquid nitrogen trap. The introduction of the precursor monomer(s) is controlled by manual or automatic valves whilst the vessel pressure is monitored. Power from a generator is coupled to the system via an external driver electrode ± seen here as a coil, although there are many other designs (Clark and Dilks, 1977; O'Toole et al., 1995; Alexander and Duc, 1998; Candan et al., 1998; Voronin et al., 2006). In order to match the impedance of the generator to that of the plasma, a matching network incorporating variable capacitors is required. In the coil configuration, the coil itself can be terminated at ground, or left as an open circuit. An important point to note is that the excitation wire passes over and around the chamber such that for RF at 13.56 MHz oscillating plasma sheaths sustain the plasma. Many researchers utilizing this type of system chose to place the non-conducting polymer substrate on the vessel floor such that any coupling of the RF through the vessel wall and substrate may induce large self-bias potentials on the substrate itself leading to energetic ions bombarding the substrate. Substrates mounted on a platform in the centre of the vessel can also experience this self-bias and energetic ion bombardment; however, these potentials are usually much smaller. The matching network, cables and radiative losses of RF power result in poor power coupling efficiencies such that in a typical coil-wound system (Fig. 1.1) as little as 20±50% of the power on the dial is actually transmitted to the plasma (Barton et al., 1999). As the quantity of this power supplied increases, the coupling efficiency decreases before reaching a maximum, saturation level. The level at which this saturation occurs is dependent on many factors including the geometry and volume of the reaction chamber, the operating pressure and the choice of gas and/or monomer(s). Since no two experimental systems are exactly the same and small differences in design can lead to large changes in power coupling efficiency, it is impossible to directly correlate processes and process parameters between systems. Attempts have been made to eliminate such variations and, for research into plasma treatment and etching, the Gaseous Electronics Conference (GEC) RF reference cell was developed (Hargis et al., 1994). However, comparative studies with these systems have highlighted the difficulty of equivalent system construction and diagnosis (Sobolewski, 1995; Graham et al., 1999). Capacitively coupled systems fitted with internal electrodes are an alternative (Hegemann et al., 2005; Vassallo et al., 2006; Hegemann et al., 2007). In such

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ß Woodhead Publishing Limited, 2011 1.1 Schematic of a typical RF (13.56 MHz) plasma system for the preparation of plasma polymer and co-polymer surfaces. Source: Figure 1, p. 421 in Haddow, D.B., MacNeil, S. and Short, R.D. (2006), `A cell therapy for chronic wounds based upon a plasma polymer delivery surface'. Plasma Processes and Polymers, 3, 419±430. Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission.

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Surface modification of biomaterials

systems one electrode is driven whilst the second, on which substrates to be coated are generally placed, acts as the return. Such systems are efficient with regards to power coupling, and adjusting the distance between the electrodes offers a further means of modifying the plasma parameters (Kawamura et al., 1999). Drawbacks include the potential to sputter the electrode material during processing ± a source of contamination in the resultant film and the changes to the physical make-up of the plasma when non-conducting polymer film builds up on the electrode surface (Roth, 1995).

1.4

Plasma parameters

Despite there being advantages and disadvantages to either method of RF power coupling, the majority of systems used for research are variations on the first with an external driven coil. The key parameters (cited in the literature) for the configurations of this type are power, as read from a dial on an external meter located on the generator, whether this power is applied continuously or pulsed, the monomer flow rate and the related parameter of the monomer pressure, and the reactor geometry. Each of these parameters affects the intrinsic parameters of the plasma and, as a result, the plasma polymer film formed and the following sections briefly discuss these parameters. For a more detailed description of technological plasmas the reader is directed to a number of excellent texts on the subject in the suggested reading section at the end of this chapter.

1.4.1

Monomer(s) flow rate

One important parameter in plasma processing is the flow rate () of the monomer gas into the system which sets the rate at which fresh monomer replaces fragmented species. This flow rate is related to the number of molecules in the vessel (N) and their residence time () by: N 1:1  Here  is expressed as the number of molecules per second but it is generally quoted and measured in standard cubic centimetres per minute (sccm). Typically monomer flow rates of the order of 1 to 20 sccm are used resulting in discharge pressures in the mTorr to 10s of mTorr range (Candan et al., 1998; Voronin et al., 2007b). For a given pumping speed of the vacuum system used to evacuate the reactor vessel,  will determine the pressure of operation. Hence, in a typical system at 10 mTorr, (N) is 3:3  1020 mÿ3 and so, with a volume of 1  10ÿ3 mÿ3 this equates to 3:3  1017 molecules in the reactor vessel such that a flow rate of 5 sccm, equivalent to 2:64  1018 molecules per second, results in a resident time () of 125 ms. For a greater pressure of 50 mTorr the resident time would equate to 0.6 s at the same flow rate whilst for systems operating with low flow rates (e.g. ˆ

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1 sccm) then () equates to 3 seconds. So we can note that the typical residence times of precursor monomer(s) are in the range of milliseconds to a few seconds.

1.4.2

Plasma power

Power is an important factor which determines such properties as the state of the plasma, its density, ion energies and the degree of monomer fragmentation (Panchalingam et al., 1993; Savage et al., 1993; Waldman et al., 1995; O'Toole et al., 1996; Coulson et al., 2000; Mota et al., 2002, Mitu et al., 2003). The following discussion focuses on RF power coupling (both continuous and pulsed) since, for the purposes of generating plasma polymer coatings for application on biomaterial and medical device surfaces, this is the method of choice. The retention of functionality from monomer(s), through the plasma to polymeric film was perhaps first described by Richard Ward in the early 1980s at the University of Durham (Ward, 1985). Indeed it was Ward who was probably first to observe the inverse relationship between plasma power and functional group retention for a series of carboxylic acids. Many studies followed in the late 1980s and early 1990s focussing on plasma power and the retention of functional groups (Savage et al., 1993; Ward and Short, 1993; Ryan et al., 1996; Alexander and Duc, 1998). For example, Ward and Short, in a study using a series of methacrylate monomer precursors, showed an inverse relationship between power and functional group retention. The explanation given being that reduced power coupled to the plasma results in reduced fragmentation of the monomer. Subsequent studies, both with other compounds and pulsed plasma, confirmed this observation and it is now generally accepted that highly functionalized films, closely resembling their analogue(s) can be obtained from the use of low power (Haddow et al., 2000a; Fraser et al., 2002). However, as shown by Ward, Short and others, perhaps a more meaningful parameter than applied power alone is the power/flow rate ratio (P=) adapted from the ratio (P=M) first described by Yasuda (Yasuda, 1985; Ward and Short, 1995; Candan et al., 1998; Chen and Yang, 1999; Oran et al., 2005). Generally, it will be seen that an increase in power and thus an increase in (P=), can lead to a reduction in the functional group retention from the monomer precursor(s). Figure 1.2 (adapted from Ward and Short, 1995) illustrates this point for plasma polymer films prepared from three methacrylate monomers. As (P=) was increased the functional group retention, here determined by the % retention of carboxylate, decreased. This (P=) ratio also gives a measure of the average energy absorbed per atom or molecule in the reactor system and thus (based upon ) establishes whether the monomer, by the action of dissociating electrons, is consumed or replenishes the fragmented gas: the former regime is known as the monomer deficient regime and the latter the power deficient regime (Yasuda, 1985). It will be seen that control of functional group density in the resultant coating is important in mediating the biological-biomaterial

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1.2 The effect of P= on functional group retention in plasma polymer films prepared from a selection of methacrylate monomer precursors (after Ward and Short, 1995). Reprinted from Polymer, 36, Ward, A.J. and Short, R.D. A t.o.f.s.i.m.s. and x.p.s. investigation of the structure of plasma polymers prepared from the methacrylate series of monomers. 2. The influence of the W/ F parameter on structural and functional-group retention, 3439±3450. Copyright 2009, with permission from Elsevier.

interaction. The average energy available per atom or molecule can be understood in terms of the parameter Emean (Bauer et al., 2005a). Emean is given by: Pabs  1:2 N with Pabs being the true, absorbed power dissipated into the plasma (not that simply read off the power dial), N the number of resident particles (monomer) and  the residence time of a particle in the plasma volume. In using continuous RF power researchers have sought to increase Pabs by increasing the input power as read on the power dial and, with the caveat identified in Section 1.3, thereby control the concentration of specific functional group retained in the coating. This technique has been extensively studied for acrylic acid discharges (O'Toole et al., 1996; Haddow et al., 2000b; Sciarratta et al., 2003; Dhayal, 2006; Voronin et al., 2007b). In general it has been found that for a molecule of acrylic acid a small value of Emean, of the order of < 20 eV/molecule, corresponds to a low degree of fragmentation and so a high degree of functional group retention. A large value of Emean of the order > 100 eV/molecule corresponds to a high degree of fragmentation and so the monomer precursor is completely dissociated and functional group retention diminishes. More recently it has become increasingly popular to pulse the plasma, typically modulating the 13.56 MHz RF in the ms to s range, which is known to Emean ˆ

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adjust the plasma gas and surface chemistry (Rinsch et al., 1996; Ryan et al., 1996; Karkari et al., 2002; Voronin et al., 2006, 2007b). Thus when the plasma is switched off there is a near instantaneous (< 1 s) collapse in the RF and DC electric fields within the plasma and the charged species diffuse to surfaces ± typically within 50±100 s for light species (M < 30 amu) and 150 s for species where M > 30 amu (Voronin et al., 2007b). The analysis of films deposited in pulsed plasma revealed that their structure depends uniquely on the average dissipated energy per source gas molecule Emean in the plasma where Emean in the pulsed regime is now given by: P P ˆ 1:3 N  where is the duty cycle, i.e. the ratio of plasma-on time to total cycle length such that ˆ ton =T, where T is the period. A Pabs of 1 W sccmÿ1 corresponds to 15.5 eV of dissipated energy per source gas molecule (Bauer et al., 2005a, 2005b). Timmons et al. undertook a comprehensive physical and chemical analysis of pulsed plasma polymerization of allyl alcohol (Rinsch et al., 1996). The results (Fig. 1.3) concluded that functional group retention in the plasma polymer film was strongly influenced by the duty cycle employed. In particular, the degree of alcohol functionality retained increased as the RF duty cycle was decreased. It should be emphasized that Equation [1.3] is only applicable if the precursor monomer(s) residence time in the plasma is much longer than the pulse cycle such that Emean can readily be tuned by controlling the duty cycle and/or the plasma power. Pulsing not only changes the applied power but also modifies plasma parameters such as plasma sheath formation and potentials across the sheath which in turn determines ion energies. Indeed for some frequency and duty regimes pulsed plasma can yield greater time-averaged plasma densities and electron temperatures than an equivalent continuous power supply. The application of both continuous and pulsed power in an RF coupled system has been modelled by Liebermann and Ashida using a global power balance (Lieberman and Ashida, 1996). Using an electropositive (argon) discharge and an electronegative (chlorine) discharge they modelled the discharge density and temperature. For an electropositive discharge under the same time-average power they noted increased electron densities (ne) for pulsed discharges as compared to equivalent continuous power ± as much as almost four times greater under certain conditions. Dhayal and Bradley showed in an acrylic acid discharge at 19.5 mTorr pressure, 500 Hz, and a 50% duty cycle that during plasma on times (ton) the electron temperature reached 4±5 eV, with an initial burst of up to 8 eV (Dhayal and Bradley, 2005). During the plasma off time (toff) the electron temperature (Te) decreased to less than 0.5 eV with ion energies of 40 eV. The plasma density (ne) was found to vary over a full duty cycle ranging from 4  1016 mÿ3 during ton and decreasing to 5  1014 mÿ3 during toff. Emean ˆ

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1.3 Functional group retention in a plasma polymer of allyl alcohol as a function of RF duty cycle. Reprinted in part with permission from Rinsch, C.L., Chen, X.L., Panchalingam, V., Eberhart, R.C., Wang, J.H. and Timmons, R.B. (1996) `Pulsed radio frequency plasma polymerization of allyl alcohol: controlled deposition of surface hydroxyl groups'. Langmuir, 12, 2995±3002. Copyright 2009 American Chemical Society.

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1.4.3

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Plasma pressure

Pressure (P) is an important parameter related to the number of molecules in the vessel (N) such that: P ˆ Nkb T

1:4

where T is the gas temperature and kb the Boltzmann constant. Increasing N, generally achieved by increasing , leads to the gas becoming collisional in that the mean free path length for neutral±neutral collisions ˆ 1=N , where  is the collisional cross section, becomes smaller than the reactor vessel dimensions (Chapman, 1980). For argon, assuming a simple geometric cross-section and billiard ball collisions  ˆ 10ÿ20 mÿ2 in a typical system at 10 mTorr, N ˆ 3:3  1020 giving ˆ 30 cm similar to the length of many vessels used. Given the typical pressure range stated in Section 1.4.1 the plasma used for polymer film formation are marginally collisional (Alexander and Duc, 1998; Voronin et al., 2006). These neutral±neutral interactions are essentially elastic scattering; however, within the polymerizing environment other important reactions can and do occur. Studies have demonstrated that ion-neutral collisions between precursor monomer units and charged oligomers, where  may be greater due to charge polarization, occur such that for these interactions is very much smaller that the vessel dimensions (Shepsis et al., 2001; Fraser et al., 2002).

1.5

Intrinsic parameters

We now proceed to identify a number of the important intrinsic parameters of plasma. Namely, potentials in the plasma, at surfaces and at the boundary between the plasma and the space charge sheath that forms in front of these surfaces, the temperature (particle energies) and the particle densities and ion flux arriving at surfaces. In the course of this discussion it will be shown how manipulation of the extrinsic properties of power and pressure can provide a degree of control of these intrinsic parameters and thereby influence the formation of the plasma polymer film. In many of the studies previously undertaken this control has often been rudimentary with results that have been incompletely understood. However, where studies of both the physical and chemical parameters have been made concurrently, a greater, albeit debatable insight into the chemical processes responsible for plasma polymer film formation have arisen (O'Toole et al., 1995, 1996; Beck et al., 1998; Fraser et al., 2002; Barton et al., 2003, 2005; Dhayal, 2006; Voronin et al., 2006, 2007a; Swindells et al., 2007).

1.5.1

Potentials in plasma and at surfaces

In a plasma, insulating or electrically isolated surfaces will assume a potential such that the fluxes of positive (ions) and negative (electrons) charge carriers

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arriving will be equal (Lieberman and Lichtenberg, 1994). In the case of classic DC plasma columns the potential of the plasma measured relative to ground, the plasma potential (Vp), is typically a few volts above the most positive surface in the discharge (Chapman, 1980). This potential ensures that positively charged ions are accelerated out of the plasma and, on average, negatively charged ions and electrons are retarded such that the net current flow of each charged species is equal and is given by that delivered by the power supply. However, the picture is different in an RF powered system. Under collisionless conditions, where the number of collisions between ions in the plasma sheath is assumed to be negligible, the potential drop across the sheath separating the surface and the plasma determines the impact energy of ions. For an insulating surface the flow of ions and electrons across the sheath must be equal as no net current can flow to the surface. In the DC case the sheath potential drop, that is the difference between the plasma potential (Vp) and floating potential (Vf), is given by: " r# kTe 1 M ‡ ln V p ÿ Vf ˆ 1:5 e 2 2m where Te is the electron temperature, M the average mass of the ions and m the electron mass (Lieberman and Lichtenberg, 1994). On the right-hand side of the equation in the bracket the 12 is due to the pre-sheath (a region of plasma in front of the sheath) potential drop while the remainder is due to the sheath. Clearly the potential drop (Vp ÿ Vf ) and hence maximum ion energy (Eimax) are a function of Te. In RF powered plasma Vp, measured relative to a grounded surface, has two components arising from the DC and RF currents such that: Vp ˆ VDC ÿ VRF

1:6

The relative magnitudes of these potentials depend strongly on a number of factors including electrode configuration and most especially (such as seen in Fig. 1.1) the effective area of the driver electrode, i.e. the external coil, and the earthed electrode, i.e. the metal flanges (Dendy, 1995). In an RF discharge the insulating substrate potential is usually more negative than the classical floating potential (Vf) arising in a DC discharge. We term this the self-bias potential (Vsb) and this potential is significant because under collisionless conditions ions crossing the plasma-polymer sheath gain a maximum energy: Eimax ˆ Vp ÿ Vsb

1:7

This self-bias potential, in the presence of RF sheath potentials and hence the maximum energy, may be expressed as: " r#   kTe 1 M kTe eVrf ln I0 1:8 ‡ ln Eimax ˆ Vp ÿ Vsb ˆ ‡ e 2 e kTe 2m

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where I0 is the modified Bessel function of zero order and Vp is the mean local plasma potential (Annaratone and Braithwaite, 1991; Annaratone et al., 1992). Consideration of Equation [1.8] shows that as RF potentials increase, the selfbias voltage Vsb becomes more negative and further, for large RF amplitudes where Vrf >> kTe =e the second term on the right-hand side approximates to Vrf and the expression reduces to: Eimax ˆ Vp ÿ Vsb  Vrf

1:9

The potential that the ions fall through to the polymer film in this case is governed by the RF potential in the plasma and not Te. Studies of continuously driven RF plasma of acrylic acid have observed sheath potentials of 60 V where 50% is due to the RF and 50% due to the electrons (Chapman, 1980). In a pulsed RF plasma using the same acrylic acid monomer sheath, potentials of up to 100 eV were observed, 80% of this being due to RF modulation in the sheath (Voronin et al., 2006). Further, in a spectrometric investigation of the plasma polymerization of acrylic acid by Haddow et al. the ion energies were determined as a function of power (Haddow et al., 2000b). Figure 1.4 illustrates this for the m/z 73 ion (i.e. [M + H]+) relative to both a grounded, conducting surface and a self-biased, non-conducting (i.e. electrically insulated) surface. It can clearly be seen from Fig. 1.4 that ions, crossing the plasma sheath, will arrive at substrate surfaces with energies dependent on both the power supplied

1.4 Peak ion energy of the m/z 73 ion arriving at grounded and self-biased surfaces as a function of P in a continuous wave plasma of acrylic acid (after Haddow et al., 2000b). Reprinted in part with permission from Haddow, D.B., France, R.M., Short, R.D., Bradley, J.W. and Barton, D. (2000) `A mass spectrometric and ion energy study of the continuous wave plasma polymerization of acrylic acid'. Langmuir, 16, 5654±5660. Copyright 2009 American Chemical Society.

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to the plasma and the nature of the substrate material. This observation once again serves to caution researchers seeking to make comparisons between different plasma systems.

1.5.2

Temperature (particle energies)

In the various types of plasma used for surface treatments a number of species including electrons, ions, atoms, metastables and, in the case of molecular gases, fragments are present. These species are not in thermodynamic equilibrium; although the electrons have elevated temperatures, the more numerous heavy particles are at ambient temperature ± it is for this reason that these plasma are often referred to as `low temperature'. The mass of ions in a plasma are typically 104±105 times the mass of electrons; however, the electrical charge on these two species is of the same magnitude (with opposite charge). This leads to the electrons experiencing acceleration from the applied electric field 104±105 times greater than ions and so the kinetic electron temperature (Te) has values of a few electron volts (eV) ± sufficient to break chemical bonds. Since the lighter electrons are much more mobile than the ions which, in relative terms, are essentially static in the plasma bulk (but not in the sheaths), they can respond to the RF field and transfer much of their energy in inelastic collisions with the monomer molecules. Even if the energy transferred is insufficient to result in fragmentation, or ionization, the compound may be raised to an excited state. One route by which the excited state may `relax' is by the emission of electromagnetic radiation and plasmas can be rich sources of high-energy (VUV) photons. Indeed, the energy of this VUV emission is sufficient to also effect chemical reactions ± particularly in polymers. Some of this electromagnetic radiation is in the visible region, hence the glow of the discharge.

1.5.3

Plasma density and ion flux

The degree of ionization in `low temperature' plasma is typically small with values of 10ÿ4 to 10ÿ6 being common. In practical terms in systems with internal electrodes this equates to only 1 charged particle per 10 000 to 100 000 neutrals which yields charged particle densities in the order of 1015 to 1016 mÿ3. In systems such as those illustrated in Fig. 1.1, however, the degree of ionization is often increased as much as two-fold with corresponding charged particle densities of 1016 to 1018 mÿ3. In almost all plasma environments the number density of positive and negative species are nearly equal. In the case of noble gas plasmas in the bulk of the plasma the number of electrons will be the equal to the number of ions, i.e. ni ˆ ne since the number of negatively charge ions is considered to be negligible. If negative ions are present such as in the presence of oxygen ions in an acrylic acid plasma or fluoride ions in a fluorocarbon

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ÿ plasma this quasi-neutrality is represented thus: n‡ i ˆ ne ‡ ni (Lieberman and Lichtenberg, 1994; Swindells et al., 2007). The plasma density is determined by the power input into the plasma minus the power lost as particles with energy leaving the boundaries and the energy being radiated away. In most RF plasma polymer discharges this plasma density is peaked at the centre and falls by less than one order of magnitude to the wall sheaths. In the wall sheaths or at a substrate the electron density falls to negligible levels faster than the ion density. At the boundary of the sheath and the plasma the ions require a minimum energy to allow the sheath to be stable. This condition, called the Bohm sheath criterion, sets the ion flux to the sheath edge. For a DC plasma or slowly varying fields in a plasma of electron density (ne) and electron temperature (Te) in the bulk, the flux of ions to a wall or substrate (ÿi ) is given by:   r 1 kTe ne ÿi ˆ exp ÿ 1:10 Mi 2

where Mi is the ionic mass which, in a multi-component plasma, is replaced by an effective mass. At 10 mTorr and 10 W input power, where typically ne  1±5  1015 mÿ3 with kTe/e  1±3 eV, ion fluxes to the substrate are about 6±12  1018 mÿ2 sÿ1 (Voronin et al., 2006). These values compare well with measurements of ion fluxes in other systems where the driver coil is external to the vessel (Beck et al., 1998; Barton et al., 1999).

1.6

Potential biomaterial applications

In addressing the application of plasma polymerization as a method of preparing biomaterial surfaces the following identifies and summarizes the perceived advantages this surface engineering technique offers: 1. Ultra thin films that do not effect bulk material. Plasma polymer coatings have thicknesses that are generally of the order of a few tens of nanometres (although the technology is capable of much greater where desired). As such the coatings have no effect on the underlying bulk material properties such as mechanical strength, flexibility, transparency, etc., that determined the initial material selection (Yasuda, 1985). 2. Conformal, uniform, pin-hole free. Plasma polymer coatings can be applied to complex geometric forms such as tissue culture plates, tubing, synthetic heart valves and porous polymer scaffolds (Tran et al., 1999; Carlisle et al., 2000; Sefton et al., 2001; Wang et al., 2005; Barry et al., 2006). 3. Can be insoluble or, where required, can be made to be soluble. Key to the process of plasma polymer film formation is the control of film chemistry or cross-link density ± properties that can be manipulated through control of plasma processing parameters such as P=. Where desired, this chemistry

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4.

5.

6.

7.

8.

Surface modification of biomaterials can be highly insoluble such as films formed from fluoro-containing monomer precursors. An example where film solubility has proven pivotal is in the development of MySkinTM, a skin cell therapy which is discussed in greater detail in Section 1.6.3 (Haddow et al., 2006). Sterile and sterilizable. A requirement of biomedical and medical devices is that they be free of any potentially harmful pathogens. The twin techniques of plasma polymerization and plasma surface treatment have both shown application as processes of sterilization (Holy et al., 2001; Haddow et al., 2003). Where additional manufacturing steps are required prior to clinical use plasma polymer coatings can withstand further sterilization by internationally accepted methods such as UV irradiation (Haddow et al., 2006). Do not necessarily age. Suppliers of biomaterials and medical devices are required to demonstrate proof of product stability or `shelf-life'. Under suitable storage conditions many plasma polymer coatings have shown resistance to adverse environmental conditions liable to cause degradation or `aging' (Whittle et al., 2000). Process readily integrated into current Good Manufacturing Processes (cGMP). The technology identified in Section 1.3 can be up-scaled to facilitate both batch and continuous manufacturing processes operated under cGMP. Medicinal products based on plasma polymer coatings have received licence to market from international regulatory bodies. As identified previously, there are a range of processing parameters that can be optimized with respect to the coating's physicochemical properties such as functional group retention and film thickness. Identifying the plasma conditions that lead to retention of functional group is perhaps the most significant development in the field since the 1980s; it provides films displaying predominately (if not exclusively) a single functionality and these can be used directly, or as a platform for assembling further chemistry or biomolecules (Kingshott et al., 2002; Siow et al., 2006). Ability to spatially pattern (controlled deposition of sub-millimetre features). Cells of different types co-cultured present both an opportunity and a challenge. On one hand co-culture has been shown to improve culture protocol with enhanced maintenance of cell phenotype; however, one cell may readily overgrow the other. The deposition of patterned plasma polymer coatings such that different chemistries are present on the one surface offers a potential route to overcome this problem (Bullett et al., 2001). A further advantage of plasma in the provision of enhanced culture surfaces is the ability to readily pattern the deposited layers, so different chemistries may be deposited on the one surface whereby each may be tailored to a specific cell type (Phillips et al., 2001).

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1.6.1

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Interaction at the biological±biomaterial interface

As noted the biological±biomaterial interaction is mediated at the surface. As a consequence it is the surfaces of these materials that are critically important in determining both the biological responses to the biomaterial and the biomaterial responses to the physiological environment (Hubbell, 1995; Babensee et al., 1998; Ratner and Bryant, 2004). These events are intended to isolate the implant (biomaterial) and subsequently heal the wound site. Indeed the resultant changes to the biomaterial surface brought about by the physiological environment, in turn, modify the biological response resulting in a complex cascade of events (Anderson et al., 2008). The primary goal, then, of biomaterials research is to design surfaces that either do not initiate a response or elicit one that is perceived to be desired. In order to tailor these biological response(s) researchers have sought to modify the physicochemical characteristics of the biomaterial such as surface energy, charge and chemistry. A number of methods have been employed, i.e. plasma treatment with inert gases and reactive gases and ion implantation (Klee and Hocker, 1999; Chu et al., 2002; McKenzie et al., 2004). Whilst these methods have found application, the focus of this chapter is the use of plasma polymer coatings which, by selection of precursor monomer(s) and processing, present a surface with the (perceived) desired physicochemical characteristics as determined by conventional wisdom. Certain features of plasma polymers, as identified in the above summary, are axiomatic to their exploitation as coatings for biomaterials and we do not comment upon these further. Other features such as functional and structural group retention are those that have been exploited and it is these that have been utilized to modify one or more physicochemical properties of a surface.

1.6.2

Plasma polymers and surface wettability (cell culture)

In 1971 Robert Baier presented a study on the role of surface energy in thrombogenesis addressing issues associated with biomaterials in contact with blood (Baier, 1972). The observations Baier made led him to propose that the most important criterion for good cardiovascular prosthesis was the surface energy of the biomaterial. An important caveat to this was that Baier qualified his observations; stating that whilst a low-critical surface energy was a practical material characteristic, this was not the only criterion to be evaluated when improving biomaterial performance. An early example of how the plasma processing conditions can control surface wettability was demonstrated for Teflon-like plasma polymer films (Favia, 1997). Using feedstock gases of C2F6 with H2 or CH4 in the glow and afterglow regions of an RF driven plasma a series of films with a range of fluorine/carbon ratios, as determined by XPS, were produced. Advancing water

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contact angle measurements on the films ranged from slightly hydrophilic to hydrophobic. This increased hydrophobicity, in good agreement with conventional polymers, correlating with increasing fluorine/carbon ratio. More recently the surface modification of NafionTM, the preferred membrane material for in situ glucose biosensors, using a range of plasma copolymer coatings was examined (Valdes et al., 2008). Plasma copolymers of tetraethylene glycol dimethyl ether (tetraglyme) and 2-hydroxyethyl methacrylate (HEMA) with varying ratios of tetraglyme : HEMA were prepared. Chemical characterisation by XPS and contact angle measurements derived a correlation between the surface wettability and hydroxyl content with this content decreasing as the ratio of HEMA in the feedstock was increased.

1.6.3

Plasma polymers and surface functional groups (cell attachment)

A successful research programme extending over several years was undertaken by groups led by Short and MacNeil at the University of Sheffield, UK, whereby lower plasma power conditions, i.e. a reduction in P=, were used to retain monomer functionality into a coating and, by a plasma copolymerization method, control functional group density (Beck et al., 1996). Note that for the preparation of chemically functionalized surfaces the use of self-assembled monolayers (SAMs) is a well-developed, alternative method (Mrksich et al., 1996; Schreiber, 2004; Ruckenstein and Li, 2005; Morgenthaler et al., 2008; Raynor et al., 2009). However, whilst SAMs have proven to be an excellent aid in studies of the biological±biomaterial interface in vitro, their application in a clinical, in vivo environment presents a number of technological challenges. For Short and MacNeil, the objective of their work was to utilize plasma polymerized coatings and, as the project developed, the subsequent delivery of skin (keratinocytes) from a single surface for wound healing. Initially the collaboration sought to establish if keratinocyte expansion on tissue culture polystyrene (TCPS) could be improved by coating with a plasma polymer in preference to collagen I (the favoured method at that time). The challenge of defining a surface for in vitro keratinocyte expansion was to establish a surface which did not promote cell differentiation but which could, in combination with appropriate culture media, maintain keratinocytes in a proliferative phenotype such that they would be capable of attaching, migrating and forming colonies. Plasma polymer coatings with different densities of acid, amine and alcohol functionality were prepared across two separate studies (France et al., 1998a, 1998b). Using collagen I coated TCPS as a `gold' standard control and uncoated TCPS as a negative control the attachment of keratinocytes on these plasma polymer coated TCPS surfaces was measured at 24 hours (Fig. 1.5) (France et al., 1998a).

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1.5 Attachment of primary keratinocytes to plasma-copolymerized surfaces at 24 hours as a percentage of attachment to collagen I. Data compiled from two studies (France et al., 1998a, 1998b).

The results from the series of acrylic acid plasma polymers were very encouraging with the measured cell number on the 2.3% acid surface at 24 hours being equivalent to that seen on collagen I. Microscopic examination of the cells on this surface and collagen I revealed cell features consistent with good attachment, spreading and colony formation. On the low acid surface, following a more extended culture period, keratinocytes proliferated and spread to form confluent sheets. This behaviour was similar to that observed on collagen I and after 7 days in culture the cells on the 2.3% acid plasma polymer were as confluent as those on the collagen I surface (Haddow et al., 1999). A more complete account of these studies, the development of a bandage (MySkinTM) for delivering skin and its clinically approval for the treatment of burns and chronic wounds, can be found elsewhere (Haddow et al., 2006). In addition, MacNeil and colleagues have continued to extend the use of plasma polymer surfaces in the expansion and delivery of other cell types including corneal and melanocyte (Eves et al., 2005; Bullock et al., 2006; Notara et al., 2007; Deshpande et al., 2009). In another study of cell-biomaterial interaction the group of Timmons examined the effect of amine functionalized plasma polymer films on neuronal cell adhesion (Harsch et al., 2000). Here, pulsed plasma films of allylamine, previously characterized by XPS and trifluoroacetic anhydride (TFAA) derivatization, were prepared using two duty cycles ± 3/45 and 3/5 (ton/toff) with the lower duty cycle films showing greater amine functional group retention (Calderon et al., 1998). At 24 hours cell dispersion was better on the higher amine functionalized (3/45) surface; however, the authors note that after two weeks culture no differences in cell morphology were detectable.

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1.6.4

Surface modification of biomaterials

Plasma polymers and spatial patterning

In developing enhanced culture surfaces a further advantage of surfaces prepared from plasma polymers is the ability to readily pattern the deposited layers (Plate I between pages 208 and 209) such that different chemistries, each tailored to a specific cell type, may be deposited on the one surface (Bullett et al., 2001). Where patterning of surfaces is desired the maintenance of pattern fidelity requires an ability to control the properties and formation of the plasma sheath (Section 1.5.1). Recently it has been shown that plasma sheath formation through masks (and indeed porous structures) influences plasma polymer formation (Zelzer et al., 2009). The relative dimensions of sheath, mask feature size and, in the case of porous three-dimensional substrates, pore diameter are critical in determining the ionic species that participate in film formation. Plasma patterning of this type can overcome an issue often encountered during the co-culture of two cell types where one cell type overgrows the other (Phillips et al., 2001).

1.6.5

Plasma polymers and surface functional group (protein binding)

As previously mentioned the surface of plasma polymer coated biomaterials and implants becomes modified with an adsorbed protein/polysaccharide layer when exposed to the biological fluid in vivo, the cell medium constituents in vitro, or by cell-deposited extracellular matrix proteins. The adsorbed protein/ polysaccharide film is fundamentally important for the success of the biomaterial because the structure and composition of the film determines the type and extent of the biological response towards the entire biomaterial. Furthermore, the nature of this adsorbed film depends on the material surface properties and so researchers have sought to adsorb specific proteins and polysaccharides to plasma polymer surfaces in a controlled manner (Lassen and Malmsten, 1997; Tang et al., 1998; Whittle et al., 2002; Zhang et al., 2003). These studies have sought to better understand: · the orientation and conformation of these adsorbed proteins and polysaccharides, · the spatial density of these and, · to what extent this prepared pseudo-biological surface can mediate the biomaterial±biological interaction. In the simplest instance, a plasma polymer coating can be used to `present' a passively adsorbed biomolecule in an optimal conformation, for subsequent cell attachment. For example, Whittle et al. (2002) have studied the passive adsorption of vitronectin, collagen and immunoglobulin G (IgG) from single protein solutions onto plasma polymer surfaces prepared from alcohol, acid,

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amine and hydrocarbon containing monomers with a range of surface energies. From this study it became apparent that there is no `magical' plasma polymer film that promotes the adsorption of active proteins and that protein adsorption is not entirely dependent on the degree of hydrophilicity of the plasma polymer surface. This work was followed up by Bullett et al. (2003), who demonstrated that the highest amount of immunologically active IgG was achieved on the more hydrophilic plasma polymer deposits. As stated by the authors in the above studies, it is apparent that protein adsorption on plasma polymers is dependent not only on wettability but on a range of other parameters including surface charge and chemistry, solubility, swelling rate and the nature of the protein of interest. A higher concentration of adsorbed protein on the plasma polymer does not necessarily correlate with the surface displaying a higher level of protein activity and these studies provide further evidence that no single physical and/or chemical property dictates biomaterial±biological interaction.

1.6.6

Plasma polymers and functional group retention (immobilization of biologically active molecules)

The `simple' adsorption of proteins and biologically active molecules to a plasma polymer surface is not the only method whereby researchers have sought to develop a pseudo-biological surface. Some of the other strategies employed include the generation of anti-fouling coatings by immobilization of polysaccharides onto amine surfaces via standard carbodiimide cross-linking chemistry or onto aldehyde surfaces through reductive amination (McArthur et al., 2000; McLean et al., 2000; Blattler et al., 2006); the attachment of enzyme to amine surfaces via a glutaraldehyde linker (Abbas et al., 2009); and the immobilization of antibacterial agents to amine surfaces based on azide/nitrene chemistry (Al-Bataineh et al., 2006). Arguably the method of choice for many researchers is the use of surface amine groups using carbodiimide chemistry (Siow et al., 2006). However, carboxylic acid groups can also be used to immobilize biomolecules through the accessible primary amines in the biomolecule. Recently, it has been reported how surface carboxylic acid groups, formed from plasma copolymerization of acrylic acid and 1,7-octadiene, may be used to covalently immobilize an anti-myc-tag (9E10) antibody. This antibody can then be used to capture (any) myc-tagged biomolecules (Walker et al., 2009). In this study the 9E10 antibody was used to immobilize the myc-tagged intercellular signalling molecule delta-like-1 (DII1). The importance of this molecule is that it regulates (inhibits) cell differentiation, and by identifying a critical surface density of DII1, it is possible to fabricate a culture surface for stem cells that permits cell expansion, without differentiation. A gradient of an amine functionalized plasma polymer surface, generated through the plasma copolymerization of allyl amine and 1,7-octadiene, has also

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1.6 Comparison of the amount of heparin bound across the length of the coverslip (represented by S atomic %, ·) and the extent of biological function as measured using the Link_TSG6 binding assay (}). The dotted line indicates the likely background reading that should be deducted from the Link_TSG6 results. Source: Figure 2, p. 1167 in Robinson, D.E., Marson, A., Short, R.D., Buttle, D.J., Day, A.J., Parry, K.L., Wiles, M., Highfield, P., Mistry, A. and Whittle, J.D. (2008) `Surface gradient of functional heparin'. Advanced Materials, 20, 1166±1169. Copyright Wiley-VCH Verlag GmbH & Co. KGaA. Reproduced with permission.

been described for study of the electrostatic adsorption of functional heparin (Robinson et al., 2008). Heparin is an important biomolecule recognized for its protein-binding capacity which is difficult to use in controlled studies of heparin±protein interactions since it does not readily bind to plastic in a functional state. This study demonstrated that whilst the amount of adsorbed heparin increased according to the surface concentration of amines in a gradient fashion (Fig. 1.6), this increased adsorption was not accompanied with a continued, corresponding rise in heparin function. The functionality of the heparin gradient was shown by using the platform to capture the Link_TSG6 macromolecule, which is known to readily bind to heparin and the activity of the Link_TSG6 determined by ELISA. The advantage of this approach is that a range of biomolecule concentrations can be screened on a single substrate surface thus reducing the amount of (potentially) very expensive biomolecules required for each assay.

1.6.7

Plasma polymers and bioactive incorporation

An alternative to covalent immobilization or adsorption of biomolecules is to encapsulate the biomolecule or drugs into a plasma polymer coating. This approach allows the drug to be delivered to the targeted site and released over an

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extended period through simple diffusion or by biodegradation of the plasma polymer. Susut and Timmons encapsulated crystals of aspirin (acetylsalicylic acid), an inflammatory analgesic, into an allyl alcohol plasma polymer by coating the crystals in a rotatable cylindrical glass plasma reactor using pulsed RF power to lower the duty cycle and so minimize drug decomposition. Power input, coating time and duty cycle were all found to affect the drug release rates with lower rates of release being observed when any one of the above parameters was increased. Another case in point is the preparation of antimicrobial coatings (Del Nobile et al., 2004; Hume et al., 2004; Jiang et al., 2004). In one study, the group of Favia utilized a silver electrode and simultaneously sputtered/deposited a plasma polymer film in an atmosphere of diethyleneglycol dimethylether (DEGDME). The poly(ethyleneoxide) (PEO)-like coating, loaded with silver ions, demonstrated enhanced inhibition to Alicyclobacillus acidoterrestris growth when compared to both an unloaded (DEGDME) plasma polymer film or an uncoated substrate of polyethylene.

1.6.8

Plasma polymers and structural retention

Plasma polymer coatings produced using N-ispopropyl acrylamide (NiPAAm) have also attracted interest due to poly(NiPAAm)'s unique reverse transition behaviour (Bullett et al., 2006). This characteristic has significant advantages in tissue engineering because the surface of poly(NiPAAm) can be used at the physiological temperature of 37 ëC as a cell adhesive layer; however, when the temperature is reduced below a critical lower value, the material uptakes water and swells. In doing so cells are repelled from the surface allowing for an entire cell sheet to be transferred to a tissue or wound site (Matsuda, 2004). Characterization of the surface chemical and mechanical properties of a plasma polymerzied NiPAAm film has recently been presented by Cheng et al. In this study, NiPAAm, polymerized in an RF discharge in a step-wise (gradual power reduction) procedure, was compared to polyNiPAAm conventionally formed through free radical polymerziation (Cheng et al., 2005). The wettability and phase transition temperature (of 31±32 ëC) was close to that of traditional polyNiPAAm and, although the plasma polymerized NiPAAm had a higher elastic moduli and decreased swelling rate, attributed to a greater degree of cross-linking in the plasma polymer film, the values were still in the range characteristic of such hydrogels.

1.6.9

Plasma polymers and function and structure retention (low biofouling plasma polymers)

Low fouling plasma polymer coatings are seen to be advantageous when minimal biological interaction with the biomaterial is desired. For instance, to improve the longevity of contact lens wear, transparent plasma polymers have

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been used as coatings to resist protein adsorption (and so, cell attachment) for several decades (Yasuda et al., 1975; Nicolson and Vogt, 2001; Weikart et al., 2001). It is well known that polymers containing ethylene oxide units (CH2± CH2±O) inhibit protein adsorption (Beyer et al., 1997; Shen et al., 2002; Zhang et al., 2003). Low or anti-biofouling polymer coatings have been achieved through methods such as grafting, covalent immobilization and chemical crosslinking (Ulbricht and Belfort, 1996; Griesser et al., 2002; Kingshott et al., 2002; Kang et al., 2008). Low biofouling plasma polymer coatings have traditionally been achieved using glymes as the precursor monomer specifically tri(ethylene glycol) and tetra(ethylene glycol) (Mar et al., 1999; Shen et al., 2002). As an alternative to long chain PEO coatings produced from traditional precursors, Wu et al. (2000) assessed the protein-resistant capability of a plasma polymer coating produced from a low molecular weight diethylene glycol vinyl ether ± EO2V. Here the lower duty cycles used in a pulsed plasma polymerization of the EO2V monomer produced films with lower water contact angles and higher C±O/C±C ratios reflecting a higher level of ethylene oxide retention with the observation made that it was these films that were most effective in resisting protein adorption. However, it was observed that the decrease in protein adsorption was not necessarily related to an increase in surface wettability ± alternatively it was proposed that strong bonding of H2O molecules to the ethylene oxide units assisted in screening the surface to proteins. This hypothesis was considered by Zhang et al. (2003) in using surface plasmon resonance (SPR) to study the adsorption of several proteins onto EO2V coatings produced by continuous wave and pulsed plasma polymerization. The films produced by pulsed plasma polymerization under the lower duty cycles displayed a lower level of crosslinking, allowing for more water to penetrate the polymer. This resulted in increased swelling after immersion in aqueous solutions and a higher level of protein resistance compared to the other coatings.

1.7

Future trends in plasma polymers

In concluding this chapter, some emerging applications of plasma polymers in biomedical engineering will be briefly discussed. The two examples given serve to illustrate the development that both fields continue to experience. The use of stem cells in the field of advanced medical treatment has garnered significant interest in recent years and plasma polymer surfaces have a role in the development of stem cell based therapies (Wells et al., 2009). Here, surface chemical gradients from a hydrophobic plasma polymerized 1,7-octadiene to a more hydrophilic plasma polymerized acrylic acid were formed on glass coverslips. Culture of E14 and R1 mouse embryonic stem cells (mES) in differing culture media was assessed on these surfaces with cell pluripotency determined by alkaline phosphatase staining. The results demonstrated that for

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these cell lines the capacity for self-renewal was maintained if cell spreading was restricted to <120 m2. The preparation of atmospheric plasma polymers is rapidly developing into a new generation of plasma polymer technology for several reasons. Atmospheric plasma have the same capabilities of low pressure plasma with additional advantages; a very large operational pressure range from mTorr to over 10s of atmospheres and the potential for the generally low temperature plasma to be miniaturized into dimensions of the order of a few microns. The scaling down of atmospheric plasma into a microplasma with a very small `footprint' enables this technology to be directly used for chemically patterning surfaces at a biologically relevant scale without the need for any masking techniques. Thus, Fig. 1.7 illustrates features (200 m in diameter) of a plasma polymer deposit of acrylic acid on a silicon wafer produced at atmospheric pressure with a microcavity discharge source. A recent report by Aizawa et al. (2007) further serves to demonstrate the flexibility of this patterning technique. Here the authors used an atmospheric plasma jet to locally deposit films of polystyrene (Fig. 1.8) that accurately reflected a pattern generated using computer aided drawing (CAD) data inputted to a motorized stage.

1.7 An XPS image of microplasma generated acrylic acid plasma polymer spots (dark) coated onto a silicon wafer (light). The image represents a surface plot of the Si 2p peak. Silicon is absent inside the areas covered by the acrylic acid plasma polymer coating, which has a spot diameter of approximately 200 m.

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1.8 Patterns of microplasma polymerized styrene deposited using a scanning microplasma jet coating system with automatic motion controller (SMPJCSAMC). Source: Figure 2, p. 218 in Aizawa, H., Makisako, T., Reddy, S. M., Terashima, K., Kurosawa, S. and Yoshimoto, M. (2007) `On-demand fabrication of microplasma-polymerized styrene films using automatic motion controller'. Journal of Photopolymer Science and Technology, 20, 215±220. Reproduced by permission of The Conference of Photopolymer Science and Technology (CPST).

In a separate approach, plasma printing technology has emerged as a tool for generating plasma polymer patterns of a very high resolution (below 100 m). This is accomplished by stamping the sample directly against the patterned microplasma source, which inhibits the diffusion of the monomer gas. Plasma polymerized amine patterns have been generated with this technology and potential commercial applications for production of flexible printed circuits are being pursued (Kreitz et al., 2005; Hinze et al., 2008; Mobius et al., 2009). Heyse et al. have recently taken plasma polymerization further by directly incorporating biomolecules in an active state into the plasma polymer deposit (Heyse et al., 2007, 2008; Ortore et al., 2008). The few examples mentioned above are only a few possibilities that atmospheric plasmas can offer. Furthering our understanding of this technology will see atmospheric plasmas become equally as important, if not more so, than existing surface engineering techniques in the very near future. In summary the use of plasma polymers in preparing modified surfaces for biomaterials has developed since the early 1970s to become a foundation technique. Rather than attempting to capture a complete history the examples above serve to highlight the role that these surfaces have played in the (continuing) evolution of our understanding of the biomaterial-biological interface. As the knowledge of plasma polymer processes has developed, so too has that of the interface and vice versa such that an almost symbiotic relationship has now emerged.

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Sources of further information and advice

Chapman, B. (1980) Glow Discharge Processes, New York, John Wiley and Sons Ltd. Dee, K. C., Puleo, D. A. & Bizios, R. (2002) An Introduction to Tissue-Biomaterial Interactions, New York, Wiley-Liss. Dendy, R. O. (1995) Plasma Physics: An Introductory Course, Cambridge, Cambridge University Press. Grill, A. (1993) Cold Plasmas in Materials Fabrication: From Fundamentals to Applications, New York, Wiley-IEEE Press. Liberman, M. A. & Lichtenberg, A. J. (1994) Principles of Plasma Discharges and Materials Processing, Chichester, John Wiley and Sons Ltd. Manos, D. M. & Flamm, D. L. (1989) Plasma Etching; An Introduction, Boston, MA, Academic Press. Ratner, B. D., Hoffman, A. S., Schoen, F. J. & Lemons, J. E. (2004) Biomaterials Science: An Introduction to Materials in Medicine, Amsterdam, Elsevier. Reece Roth, J. (1995) Industrial Plasma Engineering, Vol. 1: Principles, Bristol and Philadelphia, IOP Publishing. Williams, D. F. (1999) The Williams' Dictionary of Biomaterials, Liverpool University Press.

1.9

References

Abbas, A., Vercaigne-Mark, D., Supiot, P., Bocquet, B., Vivien, C. & Guillochon, D. (2009) Covalent attachment of trypsin on plasma polymerized allylamine. Colloids and Surfaces B: Biointerfaces, 73, 315±324. Aizawa, H., Makisako, T., Reddy, S. M., Terashima, K., Kurosawa, S. & Yoshimoto, M. (2007) On-demand fabrication of microplasma-polymerized styrene films using automatic motion controller. Journal of Photopolymer Science and Technology, 20, 215±220. Al-Bataineh, S. A., Britcher, L. G. & Griesser, H. J. (2006) XPS characterization of the surface immobilization of antibacterial furanones. Surface Science, 600, 952±962. Alexander, M. R. & Duc, T. M. (1998) The chemistry of deposits formed from acylic acid plasmas. Journal of Materials Chemistry, 8, 937±943. Anderson, J. M., Rodriguez, A. & Chang, D. T. (2008) Foreign body reaction to biomaterials. Seminars in Immunology, 20, 86±100. Annaratone, B. M. & Braithwaite, N. S. J. (1991) A comparison of a passive (filtered) and an active (driven) probe for RF plasma diagnostics. Measurement Science & Technology, 2, 795±800. Annaratone, B. M., Counsell, G. F., Kawano, H. & Allen, J. E. (1992) On the use of double probes in RF discharges. Plasma Sources Sci. Technol., 1, 232±241. Babensee, J. E., Anderson, J. M., McIntire, L. V. & Mikos, A. G. (1998) Host response to tissue engineered devices. Advanced Drug Delivery Reviews, 33, 111±139. Baier, R. E. (1972) The role of surface energy in thrombogenesis. Bulletin of the New York Academy of Medicine, 48, 257±272. Barry, J. J. A., Silva, M., Shakesheff, K. M., Howdle, S. M. & Alexander, M. R. (2005) Using plasma deposits to promote cell population of the porous interior of threedimensional poly(D,L-lactic acid) tissue-engineering scaffolds. Advanced Functional Materials, 15, 1134±1140. Barry, J. J. A., Howard, D., Shakesheff, K. M., Howdle, S. M. & Alexander, M. R. (2006)

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Using a core-sheath distribution of surface chemistry through 3D tissue engineering scaffolds to control cell ingress. Advanced Materials, 18, 1406±1410. Barton, D., Bradley, J. W., Steele, D. A. & Short, R. D. (1999) Investigating radio frequency plasmas used for the modification of polymer surfaces. Journal of Physical Chemistry B, 103, 4423±4430. Barton, D., Short, R. D., Fraser, S. & Bradley, J. W. (2003) The effect of ion energy upon plasma polymerization deposition rate for acrylic acid. Chemical Communications, 348±349. Barton, D., Shard, A. G., Short, R. D. & Bradley, J. W. (2005) The effect of positive ion energy on plasma polymerization: a comparison between acrylic and propionic acids. Journal of Physical Chemistry B, 109, 3207±3211. Bauer, M., Schwarz-Selinger, T., Kang, H. & Von Keudell, A. (2005a) Control of the plasma chemistry of a pulsed inductively coupled methane plasma. Plasma Sources Science and Technology, 14, 543±548. Bauer, M., Schwarz-Selinger, T., Jacob, W. & Von Keudell, A. (2005b) Growth precursors for a-C : H film deposition in pulsed inductively coupled methane plasmas. Journal of Applied Physics, 98. Beck, A. J., Jones, F. R. & Short, R. D. (1996) Plasma copolymerization as a route to the fabrication of new surfaces with controlled amounts of specific chemical functionality. Polymer, 37, 5537±5539. Beck, A. J., France, R. M., Leeson, A. M., Short, R. D., Goodyear, A. & Braithwaite, N. S. (1998) Ion flux and deposition rate measurements in the RF continuous wave plasma polymerisation of acrylic acid. Chemical Communications, 1221±1222. Benitez, F., Martinez, E., Galan, M., Serrat, J. & Esteve, J. (2000) Mechanical properties of plasma deposited polymer coatings. Surface & Coatings Technology, 125, 383± 387. Beyer, D., Knoll, W., Ringsdorf, H., Wang, J. H., Timmons, R. B. & Sluka, P. (1997) Reduced protein adsorption on plastics via direct plasma deposition of triethylene glycol monoallyl ether. Journal of Biomedical Materials Research, 36, 181±189. Blattler, T. M., Pasche, S., Textor, M. & Griesser, H. J. (2006) High salt stability and protein resistance of poly(L-lysine)-g-poly(ethylene glycol) copolymers covalently immobilized via aldehyde plasma polymer interlayers on inorganic and polymeric substrates. Langmuir, 22, 5760±5769. Bullett, N. A., Short, R. D., O'Leary, T., Beck, A. J., Douglas, C. W. I., Cambray-Deakin, M., Fletcher, I. W., Roberts, A. & Blomfield, C. (2001) Direct imaging of plasmapolymerized chemical micropatterns. Surface and Interface Analysis, 31, 1074± 1076. Bullett, N. A., Whittle, J. D., Short, R. D. & Douglas, C. W. I. (2003) Adsorption of immunoglobulin G to plasma-co-polymer surfaces of acrylic acid and 1,7octadiene. Journal of Materials Chemistry, 13, 1546±1553. Bullett, N. A., Talib, R. A., Short, R. D., McArthur, S. L. & Shard, A. G. (2006) Chemical and thermo-responsive characterisation of surfaces formed by plasma polymerisation of N-isopropyl acrylamide. Surface and Interface Analysis, 38, 1109±1116. Bullock, A. J., Higham, M. C. & MacNeil, S. (2006) Use of human fibroblasts in the development of a xenobiotic-free culture and delivery system for human keratinocytes. Tissue Engineering, 12, 245±255. Calderon, J. G., Harsch, A., Gross, G. W. & Timmons, R. B. (1998) Stability of plasmapolymerized allylamine films with sterilization by autoclaving. Journal of Biomedical Materials Research, 42, 597±603.

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Candan, S., Beck, A. J., O'Toole, L. & Short, R. D. (1998) Effects of `processing parameters' in plasma deposition: acrylic acid revisited. Journal of Vacuum Science & Technology A, 16, 1702±1709. Carlisle, E. S., Mariappan, M. R., Nelson, K. D., Thomes, B. E., Timmons, R. B., Constantinescu, A., Eberhart, R. C. & Bankey, P. E. (2000) Enhancing hepatocyte adhesion by pulsed plasma deposition and polyethylene glycol coupling. Tissue Engineering, 6, 45±52. Chan, C. M., Ko, T. M. & Hiraoka, H. (1996) Polymer surface modification by plasmas and photons. Surface Science Reports, 24, 3±54. Chang, J. P. & Coburn, J. W. (2003) Plasma-surface interactions. Journal of Vacuum Science & Technology A, 21, S145±S151. Chapman, B. (1980) Glow Discharge Processes, New York, John Wiley and Sons Ltd. Chen, M. & Yang, T. C. (1999) Diagnostic analysis of styrene plasma polymerization. Journal of Polymer Science Part A ± Polymer Chemistry, 37, 325±330. Cheng, X. H., Canavan, H. E., Stein, M. J., Hull, J. R., Kweskin, S. J., Wagner, M. S., Somorjai, G. A., Castner, D. G. & Ratner, B. D. (2005) Surface chemical and mechanical properties of plasma-polymerized N-isopropylacrylamide. Langmuir, 21, 7833±7841. Chu, P. K., Chen, J. Y., Wang, L. P. & Huang, N. (2002) Plasma-surface modification of biomaterials. Materials Science & Engineering R ± Reports, 36, 143±206. Clark, D. T. & Dilks, A. (1977) ESCA Applied to polymers. XV. RF Glow-discharge modification of polymers, studied by means of ESCA in terms of a direct and radiative energy-transfer model. Journal of Polymer Science, Polymer Chemistry Edition, 15, 2321±2345. Coulson, S. R., Woodward, I. S., Badyal, J. P. S., Brewer, S. A. & Willis, C. (2000) Ultralow surface energy plasma polymer films. Chemistry of Materials, 12, 2031±2038. Dai, L. M., Griesser, H. J. & Mau, A. W. H. (1997) Surface modification by plasma etching and plasma patterning. Journal of Physical Chemistry B, 101, 9548±9554. Daw, R., Brook, I. M., Devlin, A. J., Short, R. D., Cooper, E. & Leggett, G. J. (1998) A comparative study of cell attachment to self assembled monolayers and plasma polymers. Journal of Materials Chemistry, 8, 2583±2584. Del Nobile, M. A., Cannarsi, M., Altieri, C., Sinigaglia, M., Favia, P., Iacoviello, G. & D'Agostino, R. (2004) Effect of Ag-containing nano-composite active packaging system on survival of Alicyclobacillus acidoterrestris. Journal of Food Science, 69, E379±E383. Dendy, R. O. (1995) Plasma Physics: An Introductory Course, Cambridge University Press. Deshpande, P., Notara, M., Bullett, N., Daniels, J. T., Haddow, D. & MacNeil, S. (2009) Development of a surface-modified contact lens for the transfer of cultured limbal epithelial cells to the cornea for ocular surface diseases. Tissue Engineering: Part A, 15, 1±14. Dhayal, M. (2006) Control of the plasma polymerized acrylic acid film surface chemistry at low electron temperature in low pressure discharge. Journal of Vacuum Science & Technology A, 24, 1751±1755. Dhayal, M. & Bradley, J. W. (2005) Time-resolved electric probe measurements in the pulsed-plasma polymerisation of acrylic acid. Surface & Coatings Technology, 194, 167±174. Eves, P. C., Beck, A. J., Shard, A. G. & MacNeil, S. (2005) A chemically defined surface for the co-culture of melanocytes and keratinocytes. Biomaterials, 26, 7068±7081. Favia, P. (1997) Plasma Processing of Polymers, Dordrecht, Kluwer Academic Publishers.

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Flamm, D. L. (1996) Plasma chemistry, basic processes, and PECVD. In Williams, P. F. (ed.), Plasma Processing of Semiconductors: Proceedings of the NATO Advanced Study Institute. Chateau De Bonas, France, 17±28 June 1996, Kluwer Academic. Forch, R., Chifen, A. N., Bousquet, A., Khor, H. L., Jungblut, M., Chu, L. Q., Zhang, Z., Osey-Mensah, I., Sinner, E. K. & Knoll, W. (2007) Recent and expected roles of plasma-polymerized films for biomedical applications. Chemical Vapor Deposition, 13, 280±294. France, R. M., Short, R. D., Dawson, R. A. & MacNeil, S. (1998a) Attachment of human keratinocytes to plasma co-polymers of acrylic acid octa-1,7-diene and allyl amine octa-1,7-diene. Journal of Materials Chemistry, 8, 37±42. France, R. M., Short, R. D., Duval, E., Jones, F. R., Dawson, R. A. & MacNeil, S. (1998b) Plasma copolymerization of allyl alcohol 1,7-octadiene: Surface characterization and attachment of human keratinocytes. Chemistry of Materials, 10, 1176±1183. Fraser, S., Short, R. D., Barton, D. & Bradley, J. W. (2002) A multi-technique investigation of the pulsed plasma and plasma polymers of acrylic acid: Millisecond pulse regime. Journal of Physical Chemistry B, 106, 5596±5603. Graham, W. G., Mahony, C. M. O. & Steen, P. G. (1999) Electrical and optical characterisation of capacitively and inductively coupled GEC reference cells. Symposium D ± Measurement Techniques for Technological Plasmas at the 1999 EMRS Spring Meeting. Strasbourg, France, Pergamon-Elsevier Science Ltd. Griesser, H. J. & Zientek, P. (1993) Characterization of hydrophilic plasma polymercoatings on contact-lenses. Abstracts of Papers of the American Chemical Society, 206, 237±PMSE. Griesser, H. J., Chatelier, R. C., Gengenbach, T. R., Johnson, G. & Steele, J. G. (1994) Growth of human-cells on plasma polymers ± putative role of amine and amide groups. Journal of Biomaterials Science ± Polymer Edition, 5, 531±554. Griesser, H. J., Hartley, P. G., McArthur, S. L., McLean, K. M., Meagher, L. & Thissen, H. (2002) Interfacial properties and protein resistance of nano-scale polysaccharide coatings. Smart Materials & Structures, 11, 652±661. Haddow, D. B., France, R. M., Short, R. D., MacNeil, S., Dawson, R. A., Leggett, G. J. & Cooper, E. (1999) Comparison of proliferation and growth of human keratinocytes on plasma copolymers of acrylic acid/1,7-octadiene and self-assembled monolayers. Journal of Biomedical Materials Research, 47, 379±387. Haddow, D. B., Beck, A. J., France, R. M., Fraser, S., Whittle, J. D. & Short, R. D. (2000a) Mass spectral investigation of the plasma phase of a pulsed plasma of acrylic acid. Journal of Vacuum Science & Technology A ± Vacuum Surfaces and Films, 18, 3008±3011. Haddow, D. B., France, R. M., Short, R. D., Bradley, J. W. & Barton, D. (2000b) A mass spectrometric and ion energy study of the continuous wave plasma polymerization of acrylic acid. Langmuir, 16, 5654±5660. Haddow, D. B., Steele, D. A., Short, R. D., Dawson, R. A. & MacNeil, S. (2003) Plasmapolymerized surfaces for culture of human keratinocytes and transfer of cells to an in vitro wound-bed model. Journal of Biomedical Materials Research Part A, 64A, 80±87. Haddow, D. B., MacNeil, S. & Short, R. D. (2006) A cell therapy for chronic wounds based upon a plasma polymer delivery surface. Plasma Processes and Polymers, 3, 419±430. Hargis, P. J. J., Greenberg, K. E., Miller, P. A., Gerardo, J. B., Torczynski, J. R., Riley, M. E., Hebner, G. A., Roberts, J. R., J.K, O., Vanbrunt, R. J., Sobolewski, M. A., Anderson, H. M., Splichal, M. P., Mock, J. L., Bletzinger, P., A, G., Gottscho, R.

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A., Selwyn, G., Dalvie, M., Heidenreich, J. E., Butterbaugh, J. W., Brake, M. L., Passow, M. L., Pender, J., Lujan, A., Elta, M. E., Graves, D. B., Sawin, H. H., Kushner, M. J., Verdeyen, J. T., Horwath, R. & Turner, T. R. (1994) The gaseous electronics conference radio-frequency reference cell: A defined parallel-plate radio-frequency system for experimental and theoretical-studies of plasmaprocessing discharges. Review of Scientific Instruments, 65, 140±154. Harsch, A., Calderon, J., Timmons, R. B. & Gross, G. W. (2000) Pulsed plasma deposition of allylamine on polysiloxane: a stable surface for neuronal cell adhesion. Journal of Neuroscience Methods, 98, 135±144. Hegemann, D., Oehr, C. & Fischer, A. (2005) Design of functional coatings. Journal of Vacuum Science & Technology A, 23, 5±11. Hegemann, D., Hossain, M. M., Korner, E. & Balazs, D. J. (2007) Macroscopic description of plasma polymerization. Plasma Processes and Polymers, 4, 229±238. Heyse, P., Dams, R., Paulussen, S., Houthofd, K., Janssen, K., Jacobs, P. A. & Sels, B. F. (2007) Dielectric barrier discharge at atmospheric pressure as a tool to deposit versatile organic coatings at moderate power input. Plasma Processes and Polymers, 4, 145±157. Heyse, P., Roeffaers, M. B. J., Paulussen, S., Hofkens, J., Jacobs, P. A. & Sels, B. F. (2008) Protein immobilization using atmospheric-pressure dielectric-barrier discharges: A route to a straightforward manufacture of bioactive films. Plasma Processes and Polymers, 5, 186±191. Hinze, A., Klages, C. P., Zanker, A., Thomas, M., Wirth, T. & Unger, W. E. S. (2008) ToF-SIMS imaging of DBD-plasma-printed microspots on BOPP substrates. Plasma Processes and Polymers, 5, 460±470. Hiratsuka, A., Muguruma, H., Lee, K. H. & Karube, I. (2004) Organic plasma process for simple and substrate-independent surface modification of polymeric BioMEMS devices. Biosensors & Bioelectronics, 19, 1667±1672. Holy, C. E., Cheng, C., Davies, J. E. & Shoichet, M. S. (2001) Optimizing the sterilization of PLGA scaffolds for use in tissue engineering. Biomaterials, 22, 25± 31. Hubbell, J. A. (1995) Biomaterials in tissue engineering. Bio-Technology, 13, 565±576. Hume, E. B. H., Baveja, J., Muir, B., Schubert, T. L., Kumar, N., Kjelleberg, S., Griesser, H. J., Thissen, H., Read, R., Poole-Warren, L. A., Schindhelm, K. & Willcox, M. D. P. (2004) The control of Staphylococcus epidermidis biofilm formation and in vivo infection rates by covalently bound furanones. Biomaterials, 25, 5023±5030. Jiang, H. Q., Manolache, S., Wong, A. C. L. & Denes, F. S. (2004) Plasma-enhanced deposition of silver nanoparticles onto polymer and metal surfaces for the generation of antimicrobial characteristics. Journal of Applied Polymer Science, 93, 1411±1422. Kang, G. D., Cao, Y. M., Zhao, H. Y. & Yuan, Q. (2008) Preparation and characterization of crosslinked poly(ethylene glycol) diacrylate membranes with excellent antifouling and solvent-resistant properties. Journal of Membrane Science, 318, 227±232. Karkari, S. K., Backer, H., Forder, D. & Bradley, J. W. (2002) A technique for obtaining time- and energy-resolved mass spectroscopic measurements on pulsed plasmas. Measurement Science & Technology, 13, 1431±1436. Kawamura, E., Vahedi, V., Lieberman, M. A. & Birdsall, C. K. (1999) Ion energy distributions in rf sheaths; review, analysis and simulation. Plasma Sources Science & Technology, 8, R45±R64. Kingshott, P., Thissen, H. & Griesser, H. J. (2002) Effects of cloud-point grafting, chain

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length, and density of PEG layers on competitive adsorption of ocular proteins. Biomaterials, 23, 2043±2056. Klee, D. & Hocker, H. (1999) Polymers for biomedical applications: Improvement of the interface compatibility. Biomedical Applications: Polymer Blends, 149, 1±57. Kreitz, S., Penache, C., Thomas, M. & Klages, C. P. (2005) Patterned DBD treatment for area-selective metallization of polymers-plasma printing. Surface & Coatings Technology, 200, 676±679. Lassen, B. & Malmsten, M. (1997) Competitive protein adsorption at plasma polymer surfaces. Journal of Colloid and Interface Science, 186, 9±16. Lieberman, M. A. & Ashida, S. (1996) Global models of pulse-power-modulated high-density, low-pressure discharges. Plasma Sources Science and Technology, 5, 145±158. Lieberman, M. A. & Lichtenberg, A. J. (1994) Principles of Plasma Discharges & Materials Processing, Chichester, John Wiley and Sons Ltd. Liston, E. M., Martinu, L. & Wertheimer, M. R. (1993) Plasma surface modification of polymers for improved adhesion ± a critical review. Journal of Adhesion Science and Technology, 7, 1091±1127. Manos, D. M. & Flamm, D. L. (1989) Plasma Etching: An Introduction, Boston, MA, Academic Press. Mar, M. N., Ratner, B. D. & Yee, S. S. (1999) An intrinsically protein-resistant surface plasmon resonance biosensor based upon a RF-plasma-deposited thin film. Sensors and Actuators B ± Chemical, 54, 125±131. Matsuda, T. (2004) Poly(N-isopropylacrylamide)-grafted gelatin as a thermoresponsive cell-adhesive, mold-releasable material for shape-engineered tissues. Journal of Biomaterials Science ± Polymer Edition, 15, 947±955. McArthur, S. L., McLean, K. M., Kingshott, P., St John, H. A. W., Chatelier, R. C. & Griesser, H. J. (2000) Effect of polysaccharide structure on protein adsorption. Colloids and Surfaces B ± Biointerfaces, 17, 37±48. McKenzie, D. R., Newton-McGee, K., Ruch, P., Bilek, M. M. & Gan, B. K. (2004) Modification of polymers by plasma-based ion implantation for biomedical applications. Surface & Coatings Technology, 186, 239±244. McLean, K. M., Johnson, G., Chatelier, R. C., Beumer, G. J., Steele, J. G. & Griesser, H. J. (2000) Method of immobilization of carboxymethyl-dextran affects resistance to tissue and cell colonization. Colloids and Surfaces B ± Biointerfaces, 18, 221±234. Mitu, B., Bauer-Gogonea, S., Leonhartsberger, H., Lindner, M., Bauer, S. & Dinescu, G. (2003) Plasma-deposited parylene-like thin films: process and material properties. Surface & Coatings Technology, 174, 124±130. Mobius, A., Elbick, D., Weidlich, E. R., Feldmann, K., Schuessler, F., Borris, J., Thomas, M., Zanker, A. & Klages, C. P. (2009) Plasma-printing and galvanic metallization hand in hand ± a new technology for the cost-efficient manufacture of flexible printed circuits. Electrochimica Acta, 54, 2473±2477. Morgenthaler, S., Zink, C. & Spencer, N. D. (2008) Surface-chemical and -morphological gradients. Soft Matter, 4, 419±434. Mota, R. P., Bigansolli, A. R., Antunes, E. F., Rangel, E. C., Da Cruz, N. C., Honda, R. Y., Algatti, M. A., Aramaki, E. A. & Kayama, M. E. (2002) Optical and structural properties of PEO-like plasma polymers. Molecular Crystals and Liquid Crystals, 374, 415±420. Mrksich, M., Chen, C. S., Xia, Y. N., Dike, L. E., Ingber, D. E. & Whitesides, G. M. (1996) Controlling cell attachment on contoured surfaces with self-assembled monolayers of alkanethiolates on gold. Proceedings of the National Academy of Sciences of the United States of America, 93, 10775±10778.

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Nicolson, P. C. & Vogt, J. (2001) Soft contact lens polymers: an evolution. Biomaterials, 22, 3273±3283. Notara, M., Bullett, N. A., Deshpande, P., Haddow, D. B., MacNeil, S. & Daniels, J. T. (2007) Plasma polymer coated surfaces for serum-free culture of limbal epithelium for ocular surface disease. Journal of Materials Science-Materials in Medicine, 18, 329±338. O'Toole, L., Beck, A. J., Ameen, A. P., Jones, F. R. & Short, R. D. (1995) Radiofrequency-induced plasma polymerisation of propenoic acid and propanoic acid. Journal of the Chemical Society-Faraday Transactions, 91, 3907±3912. O'Toole, L., Beck, A. J. & Short, R. D. (1996) Characterization of plasma polymers of acrylic acid and propanoic acid. Macromolecules, 29, 5172±5177. Oran, U., Swaraj, S., Friedrich, J. F. & Unger, W. E. S. (2005) Surface analysis of plasma deposited polymer films, 5 ± ToF-SSIMS characterization of plasma deposited allyl alcohol films. Plasma Processes and Polymers, 2, 563±571. Ortore, M. G., Sinibaldi, R., Heyse, P., Paulussen, S., Bernstorff, S., Sels, B., Mariani, P., Rustichelli, F. & Spinozzi, F. (2008) Grazing-incidence small-angle X-ray scattering from alkaline phosphatase immobilized in atmospheric plasmapolymer coatings. Applied Surface Science, 254, 5557±5563. Panchalingam, V., Poon, B., Huo, H. H., Savage, C. R., Timmons, R. B. & Eberhart, R. C. (1993) Molecular-surface tailoring of biomaterials via pulsed RF plasma discharges. Journal of Biomaterials Science ± Polymer Edition, 5, 131±145. Phillips, J., Gawkrodger, D. J., Caddy, C. M., Hedley, S., Dawson, R. A., Smith-Thomas, L., Freedlander, E. & MacNeil, S. (2001) Keratinocytes suppress TRP-1 expression and reduce cell number of co-cultured melanocytes ± implications for crafting of patients with vitiligo. Pigment Cell Research, 14, 116±125. Ratner, B. D. & Bryant, S. J. (2004) Biomaterials: Where we have been and where we are going. Annual Review of Biomedical Engineering, 6, 41±75. Raynor, J. E., Capadona, J. R., Collard, D. M., Petrie, T. A. & Garcia, A. J. (2009) Polymer brushes and self-assembled monolayers: versatile platforms to control cell adhesion to biomaterials (Review). Biointerphases, 4, FA3±FA16. Rinsch, C. L., Chen, X. L., Panchalingam, V., Eberhart, R. C., Wang, J. H. & Timmons, R. B. (1996) Pulsed radio frequency plasma polymerization of allyl alcohol: controlled deposition of surface hydroxyl groups. Langmuir, 12, 2995±3002. Robinson, D. E., Marson, A., Short, R. D., Buttle, D. J., Day, A. J., Parry, K. L., Wiles, M., Highfield, P., Mistry, A. & Whittle, J. D. (2008) Surface gradient of functional heparin. Advanced Materials, 20, 1166±1169. Roth, J. R. (1995) Industrial Plasma Engineering, Vol. 1: Principles, Bristol and Philadelphia, IOP Publishing. Ruckenstein, E. & Li, Z. F. (2005) Surface modification and functionalization through the self-assembled monolayer and graft polymerization. Advances in Colloid and Interface Science, 113, 43±63. Ryan, M. E., Hynes, A. M. & Badyal, J. P. S. (1996) Pulsed plasma polymerization of maleic anhydride. Chemistry of Materials, 8, 37±42. Savage, C. R., Timmons, R. B. & Lin, J. W. (1993) Spectroscopic characterization of films obtained in pulsed radiofrequency plasma discharges of fluorocarbon monomers. Advances in Chemistry Series, 745±768. Schreiber, F. (2004) Self-assembled monolayers: from `simple' model systems to biofunctionalized interfaces. Journal of Physics-Condensed Matter, 16, R881± R900. Sciarratta, V., Vohrer, U., Hegemann, D., Muller, M. & Oehr, C. (2003) Plasma

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functionalization of polypropylene with acrylic acid. Surface & Coatings Technology, 174, 805±810. Sefton, M. V., Sawyer, A., Gorbet, M., Black, J. P., Cheng, E., Gemmell, C. & PottingerCooper, E. (2001) Does surface chemistry affect thrombogenicity of surface modified polymers? Journal of Biomedical Materials Research, 55, 447±459. Shen, M. C., Martinson, L., Wagner, M. S., Castner, D. G., Ratner, B. D. & Horbett, T. A. (2002) PEO-like plasma polymerized tetraglyme surface interactions with leukocytes and proteins: in vitro and in vivo studies. Journal of Biomaterials Science ± Polymer Edition, 13, 367±390. Shepsis, L. V., Pedrow, P. D., Mahalingam, R. & Osman, M. A. (2001) Modeling and experimental comparison of pulsed plasma deposition of aniline. Thin Solid Films, 385, 11±21. Sibarani, J., Takai, M. & Ishihara, K. (2007) Surface modification on microfluidic devices with 2-methacryloyloxyethyl phosphorylcholine polymers for reducing unfavorable protein adsorption. Colloids and Surfaces B ± Biointerfaces, 54, 88± 93. Siow, K. S., Britcher, L., Kumar, S. & Griesser, H. J. (2006) Plasma methods for the generation of chemically reactive surfaces for biomolecule immobilization and cell colonization ± a review. Plasma Processes and Polymers, 3, 392±418. Sobolewski, M. A. (1995) Current and voltage measurements in the gaseous electronics conference RF reference cell. Journal of Research of the National Institute of Standards and Technology, 100, 341±351. Swindells, I., Voronin, S. A., Fotea, C., Alexander, M. R. & Bradley, J. W. (2007) Detection of negative molecular ions in acrylic acid plasma: Some implications for polymerization mechanisms. Journal of Physical Chemistry B, 111, 8720±8722. Tang, L. P., Wu, Y. L. & Timmons, R. B. (1998) Fibrinogen adsorption and host tissue responses to plasma functionalized surfaces. Journal of Biomedical Materials Research, 42, 156±163. Tran, H. S., Puc, M. M., Hewitt, C. W., Soll, D. B., Marra, S. W., Simonetti, V. A., Cilley, J. H. & Delrossi, A. J. (1999) Diamond-like carbon coating and plasma or glow discharge treatment of mechanical heart valves. Journal of Investigative Surgery, 12, 133±140. Ulbricht, M. & Belfort, G. (1996) Surface modification of ultrafiltration membranes by low temperature plasma. 2. Graft polymerization onto polyacrylonitrile and polysulfone. Journal of Membrane Science, 111, 193±215. Valdes, T. I., Ciridon, W., Ratner, B. D. & Bryers, J. D. (2008) Surface modification of a perfluorinated ionomer using a glow discharge deposition method to control protein adsorption. Biomaterials, 29, 1356±1366. Vassallo, E., Cremona, A., Laguardia, L. & Mesto, E. (2006) Preparation of plasmapolymerized SiOx-like thin films from a mixture of hexamethyldisiloxane and oxygen to improve the corrosion behaviour. Surface & Coatings Technology, 200, 3035±3040. Villermaux, F., Tabrizian, M., Yahia, L., Czeremuszkin, G. & Piron, D. L. (1996) Corrosion resistance improvement of NiTi osteosynthesis staples by plasma polymerized tetrafluoroethylene coating. Bio-Medical Materials and Engineering, 6, 241±254. Voronin, S. A., Alexander, M. R. & Bradley, J. W. (2006) Time-resolved mass and energy spectral investigation of a pulsed polymerising plasma struck in acrylic acid. Surface & Coatings Technology, 201, 768±775. Voronin, S. A., Bradley, J. W., Fotea, C., Zelzer, M. & Alexander, M. R. (2007a)

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Characterization of thin-film deposition in a pulsed acrylic acid polymerizing discharge. Journal of Vacuum Science & Technology A, 25, 1093±1097. Voronin, S. A., Zelzer, M., Fotea, C., Alexander, M. R. & Bradley, J. W. (2007b) Pulsed and continuous wave acrylic acid radio frequency plasma deposits: Plasma and surface chemistry. Journal of Physical Chemistry B, 111, 3419±3429. Waldman, D. A., Ziu, Y. L. & Netravali, A. N. (1995) Ethylene ammonia plasma polymer deposition for controlled adhesion of graphite fibers to PEEK. 1. Characterization of plasma formed polymers. Journal of Adhesion Science and Technology, 9, 1475± 1503. Walker, R. A., Cunliffe, V. T., Whittle, J. D., Steele, D. A. & Short, R. D. (2009) Submillimeter-scale surface gradients of immobilized protein ligands. Langmuir, 25, 4243±4246. Wang, J., Pan, C. J., Huang, N., Sun, H., Yang, P., Leng, Y. X., Chen, J. Y., Wan, G. J. & Chu, P. K. (2005) Surface characterization and blood compatibility of poly(ethylene terephthalate) modified by plasma surface grafting. Surface & Coatings Technology, 196, 307±311. Ward, A. J. & Short, R. D. (1993) A time-of-flight secondary-ion mass-spectrometry and X-ray photoelectron-spectroscopy investigation of the structure of plasma polymers prepared from the methacrylate series of monomers. Polymer, 34, 4179±4185. Ward, A. J. & Short, R. D. (1995) A t.o.f.s.i.m.s. and x.p.s. investigation of the structure of plasma polymers prepared from the methacrylate series of monomers. 2. The influence of the W/F parameter on structural and functional-group retention. Polymer, 36, 3439±3450. Ward, R. J. (1985) An investigation of the plasma polymerisation of organic compounds and its relationship with surface photopolymerisation. Department of Chemistry. Durham, University of Durham. Weikart, C. M., Matsuzawa, Y., Winterton, L. & Yasuda, H. K. (2001) Evaluation of plasma polymer-coated contact lenses by electrochemical impedance spectroscopy. Journal of Biomedical Materials Research, 54, 597±607. Wells, N., Baxter, M. A., Turnbull, J. E., Murray, P., Edgar, D., Parry, K. L., Steele, D. A. & Short, R. D. (2009) The geometric control of E14 and R1 mouse embryonic stem cell pluripotency by plasma polymer surface chemical gradients. Biomaterials, 30, 1066±1070. Whittle, J. D., Short, R. D., Douglas, C. W. I. & Davies, J. (2000) Differences in the aging of allyl alcohol, acrylic acid, allylamine, and octa-1,7-diene plasma polymers as studied by X-ray photoelectron spectroscopy. Chemistry of Materials, 12, 2664± 2671. Whittle, J. D., Bullett, N. A., Short, R. D., Douglas, C. W. I., Hollander, A. P. & Davies, J. (2002) Adsorption of vitronectin, collagen and immunoglobulin-G to plasma polymer surfaces by enzyme linked immunosorbent assay (ELISA). Journal of Materials Chemistry, 12, 2726±2732. Wohlrab, C. & Hofer, M. (1995) Application of plasma polymerization on ophthalmic lenses ± equipment and processes. Optical Engineering, 34, 2712±2718. Wu, Y. L. J., Timmons, R. B., Jen, J. S. & Molock, F. E. (2000) Non-fouling surfaces produced by gas phase pulsed plasma polymerization of an ultra low molecular weight ethylene oxide containing monomer. Colloids and Surfaces B ± Biointerfaces, 18, 235±248. Yasuda, H. (1985) Plasma Polymerization, London, Academic Press. Yasuda, H., Bumgarner, M. O., Marsh, H. C., Yamanashi, B. S., Devito, D. P., Wolbarsht, M. L., Reed, J. W., Bessler, M., Landers III, M. B., Hercules, D. M. & Carver, J.

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(1975) Ultrathin coating by plasma polymerization applied to corneal contact lens. Journal of Biomedical Materials Research, 9, 629±643. Zelzer, M., Majani, R., Bradley, J. W., Rose, F., Davies, M. C. & Alexander, M. R. (2008) Investigation of cell-surface interactions using chemical gradients formed from plasma polymers. Biomaterials, 29, 172±184. Zelzer, M., Scurr, D., Abdullah, B., Urquhart, A. J., Gadegaard, N., Bradley, J. W. & Alexander, M. R. (2009) Influence of the plasma sheath on plasma polymer deposition in advance of a mask and down pores. Journal of Physical Chemistry B, 113, 8487±8494. Zhang, Z., Menges, B., Timmons, R. B., Knoll, W. & Forch, R. (2003) Surface plasmon resonance studies of protein binding on plasma polymerized di(ethylene glycol) monovinyl ether films. Langmuir, 19, 4765±4770.

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Plate I Attachment of rat bone cancer cells (ROS cells) to regions of a polystyrene surface that have been plasma-coated using acrylic acid. The grid (top) was used to create the pattern. Scale bar is 50 micrometres.

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Surface modification of biomaterials by covalent binding of poly(ethylene glycol) (PEG) A . R H O D E S , S . S . S A N D H U and S . J . O N I S , BioInteractions Ltd, UK

Abstract: Covalent binding represents a logical approach for modifying the surface properties of a biomaterial, allowing the bulk composition to remain unchanged, whilst surface functionality is engineered to afford a more desirable characteristic. A range of different molecules have been utilised in order to achieve this goal. Of those employed, poly(ethylene glycol) (PEG) has been widely used as a result of hydrophilicity, a high exclusion volume in water, flexibility, non-toxicity and non-immunogenicity. In this review, we describe various strategies that have been used to afford covalent binding of PEG onto a range of biomaterial surfaces. The main focus will be on the methods and functional groups employed to achieve this end. Key words: covalent binding, poly(ethylene glycol), surface modification.

2.1

Introduction

The interfacial characteristics of a biomaterial's surface play an important role in determining the biocompatibility and, ultimately, the success of such a material. Indeed, the progress of promising materials, which display ideal bulk properties, can be hindered as a result of undesirable surface properties. Covalent binding represents a logical approach for modifying the surface properties of a biomaterial, allowing the bulk composition to remain unchanged, whilst surface functionality is engineered to afford a more desirable characteristic. This strategy offers wide scope for the attachment of various functional moieties onto a range of different substrates, with the inherent advantage of chemical stability, afforded through the covalent bond. This advantageous feature, coupled with the flexibility, reproducibility and relative ease of this technique, demonstrates the suitability of covalent binding for modifying the surface of biomaterials. A wide range of different molecules have been utilised in order to covalently modify biomaterial surfaces. Of those employed, poly(ethylene glycol) (PEG) (or poly(ethylene oxide) (PEO) when the molecular weight is greater than 10 000 gmolÿ1) has been widely used (Bailey and Koleske, 1976). Covalent modification with PEG provides a useful model to discuss the various strategies

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offered by covalent binding and will therefore be used as an exemplary molecule throughout this chapter. PEG is a water soluble amphiphilic polyether that demonstrates unique properties, including hydrophilicity, a high exclusion volume in water (as a result of high conformational entropy), flexibility, non-toxicity and nonimmunogenicity (Lee et al., 1995). As a result of these unique properties, PEG effectively decreases non-specific protein binding and cellular adhesion, thereby improving the biocompatibility of biomaterials. The physical adsorption of PEG has been investigated as a means of preparing PEG-rich surfaces. However, due to the highly hydrophilic nature of PEG this approach permits displacement of the polymer in aqueous media and as such, reduces the longevity of PEG on the surface (Tseng et al., 1995). Covalent binding therefore represents a logical approach for increasing the stability and hence, lifetime, of surfaces functionalised with PEG. In this review, we describe various strategies that have been used to afford covalent binding of PEG onto a range of biomaterial surfaces. The main focus will be on the methods and functional groups employed to achieve covalent modification of biomaterials.

2.2

Principles and methods

The underlying goal of covalent binding is to achieve a stable (chemical) bond between the biomaterial surface and the functional species being attached. Broadly speaking, covalent binding can be categorised into three main groups: 1. Direct reaction between functional groups on the surface and functional groups present on the species being immobilised. 2. Pre-activation of the surface, the species being immobilised or both, to provide functional groups suitable for binding. This may encompass the use of linkers and coupling agents. 3. The use of external stimuli to promote grafting reactions, such as UV irradiation or plasma. Covalent binding also offers the possibility of pre-fabricating the species being immobilised prior to coupling, offering additional control over chain length, chain structure and functionality. Typically, the functional groups employed in covalent binding include groups such as hydroxyl (±OH), amino (±NH2), thiol (±SH), carboxyl (±COOH), aldehyde (±COH), halide (±X, where X = Cl or Br) and isocyanate (±NCO), which provide suitable sites to direct covalent bond formation. As highlighted, these groups may be modified prior to the immobilisation step in order to achieve a more efficient reaction. The chemical structure of PEG, illustrated in Fig. 2.1, highlights the terminal hydroxyl groups present at either end of the polyether chain. These hydroxyl

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2.1 Chemical structure of poly(ethylene glycol) (PEG).

groups represent suitable sites for directing covalent binding to a surface and can be used directly, or suitably modified, in order to achieve PEG immobilisation. As discussed, a range of functional moieties may be employed for directing covalent bond formation. Of these, isocyanates represent an efficient group for covalently linking PEG to surfaces.

2.2.1

Isocyanates

The suitability of isocyanates for covalently immobilising PEG onto surfaces lies in the fact that no by-products are formed during the coupling reaction. In addition, this reaction is a one-step process, whereby the isocyanate group reacts directly with the hydroxyl group(s) of PEG, forming a carbamate (urethane) bond. Isocyanates can be categorised as being either aromatic or aliphatic, with aromatic isocyanates typically displaying an increased reactivity rate. This increased reactivity rate means that it is often not necessary to employ a catalyst when carrying out a reaction with hydroxyl groups (Graham, 1987). However, when aliphatic isocyanates are utilised, tertiary amines or soluble compounds of metals, such as tin, may be used in such a role. This has the obvious drawback of catalyst removal following reaction; generally, however, this can be suitably achieved through thorough rinsing/washing following the immobilisation step. The bifunctional aliphatic isocyanate, hexamethylene diisocyanate (HMDI), has been used to covalently link PEG to a surface. Han et al. (1989) first reacted HMDI onto a poly(urethane) (PU) sheet, in the presence of toluene and stannous octoate (used as catalyst), for 1 h at 40 ëC under nitrogen. Following washing with toluene and anhydrous ether, PEG was grafted onto the free isocyanate groups of the PU-HMDI sheet through a carbamate linkage. The reaction was carried out in benzene at 40 ëC for 24 h with stannous octoate again used as the catalyst. Washing with benzene followed by ethanol and subsequent drying at 20 ëC, afforded the PU-PEG sheet illustrated in Fig. 2.2. Byun et al. (1995) utilised the aromatic toluene diisocyanate (TDI) to modify the terminal hydroxyl groups of PEG. The reaction was carried out in benzene under a nitrogen atmosphere at 60 ëC for 24 h. Following precipitation in ether,

2.2 PU-PEG sheet coupled through HMDI linker.

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2.3 TDI-modified PEG.

filtration and drying under nitrogen, the TDI-modified PEG (Fig. 2.3) was coupled onto the amine groups of a model substrate, poly(styrene-co-p-amino styrene). The coupling step was carried out in formamide under nitrogen at 25 ëC for 24 h. In this step, a carbamide (urea) linkage was formed as a result of terminal isocyanates reacting with surface amine groups. For those materials that do not display surface functional groups suitable for covalent binding, such as poly(ethylene) (PE) and poly(propylene) (PP), functional groups must first be introduced. A variety of treatments are available for achieving this end, including corona discharge, UV irradiation and plasma (Goddard and Hotchkiss, 2007). The type of functionalisation imparted can be varied, for example, by the plasma gas selected (O2, NH3, Ar, etc.) and can be used to add oxygen containing functional groups as well as amine functionalities. Therefore, plasma treatment of an `inert' surface can render it suitable for the attachment of PEG via isocyanate linkages. Due to the fact that isocyanates react readily with hydroxyl groups, it is important that the correct molar ratios of isocyanate to PEG are used to ensure that cross-linking or polymerisation does not occur. One approach for avoiding this problem is to use the mono-functionalised PEG derivative, poly(ethylene glycol) monomethyl ether (mPEG).

2.2.2

Poly(ethylene glycol) monomethyl ether (mPEG)

The mono-functionalised derivative, mPEG (Fig. 2.4), allows the beneficial properties of PEG to be exploited, whilst eliminating the chance of crosslinking. By employing mPEG in place of PEG, unwanted side reactions can be reduced, thereby increasing the coupling efficiency. Kingshott et al. (2002) have utilised mPEG functionalised with a terminal aldehyde group to graft PEG onto aminated surfaces, i.e. surfaces displaying amine groups. Grafting was achieved via reductive amination, using sodium cyanoborohydride (NaCNBH3) as the reducing agent for the intermediate Schiff base (imine) linkage. The effect of temperature and salt (K2SO4) content of the PEG grafting solution were investigated, which highlighted that optimised PEG

2.4 Chemical structure of poly(ethylene glycol) monomethyl ether (mPEG).

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grafting is achieved at the `cloud point' (elevated temperature and salt content), when PEG is only marginally solvated. Reductive amination has also been used to covalent link mPEG onto poly(vinylidene fluoride) (PVDF) films, with the goal of producing a more biocompatible surface. Ademovic et al. (2002) first created carboxyl groups on the surface through plasma induced graft polymerisation of acrylic acid, to which poly(ethyleneimine) (PEI) was covalently coupled. Aldehyde terminated mPEG was then grafted onto surface amine groups by reductive amination, again using NaCNBH3 as the reducing agent. Interestingly, the grafting was also carried out under marginal solvation conditions. As well as monofunctional PEGs, a range of bifunctional derivatives are also available and have been readily employed to modify the surfaces of biomaterials.

2.2.3

Bifunctional PEGs

Bifunctional PEGs have been extensively used, particularly as linkers, for the immobilisation of bioactive molecules onto surfaces, as they provide flexibility, hydrophilicity and biocompatibility (Chen et al., 2005). In addition, a variety of functionalities can be introduced at the peripheries of the PEG chain, thus allowing for a range of different chemistries to be used. Janissen et al. (2009) have utilised a heterobifunctional PEG, with a terminal N-hydroxysuccinimide (NHS) activated carboxylic acid group (Fig. 2.5), to graft PEG onto aminated glass slides in the presence of triethylamine (NEt3) and chloroform. NHS activation increases the electrophilicity of the terminal carboxylic acid group, increasing its susceptibility to nucleophilic attack and therefore, helps to promote amide bond formation between surface amine groups and PEG. Having covalently linked the heterobifunctional PEG to the surface, the terminal carboxyl group was then used to immobilise bioactive molecules using 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride (EDC) as the coupling agent. The EDC coupling method has also been utilised by Piehler et al. (1996) to attach a homobifunctional PEG species (Fig. 2.6) onto glass-type surfaces. Initially, the surface was activated by silanisation with an amino-functionalised silane, namely 4-aminobutyldimethylmonomethoxysilane (ABDMS), to afford

2.5 Chemical structure of heterobifunctional PEG possessing a terminal NHS group.

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2.6 Chemical structure of homobifunctional diamino-PEG.

2.7 Surface activation using ABDMS and succininc anhydride.

surface amine groups. Subsequent ring-opening of succinic anhydride yielded surface carboxyl groups, which provided suitable sites for coupling the homobifunctional diamino-PEG species onto the surface (Fig. 2.7). Coupling via an amide bond was achieved either directly, using excess EDC at pH 3±4 for 12 h, or through pre-activation of the carboxyl groups with N,N'diisopropylcarbodiimide (DIC) and NHS in the presence of dimethylformamide (DMF) for 4 h (followed by incubation in the polymer solution for 18 h). Although various functional groups can be introduced at the terminal ends of PEG, it is often necessary to employ agents, such as EDC, to promote coupling reactions and therefore increase their efficiency.

2.2.4

Coupling compounds

EDC, DIC and N,N0 -dicyclohexylcarbodiimide (DCC) (Fig. 2.8) can all be used to `activate' carboxylic acids, thereby increasing their reactivity towards amines and hydroxyl groups. The simultaneous use of NHS in such reactions serves to

2.8 EDC, DIC and DCC.

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2.9 1,10 -carbonyldiimidazole (CDI).

decrease side-reactions by stabilising the `activated' intermediate formed and, therefore, can be beneficial for improving reaction yields. The disadvantage of using DCC and DIC, however, is that the urea by-product (N,N0 -dicyclohexylurea and N,N0 -diisopropylurea, respectively) is often difficult to remove. However, when coupling reactions are performed on a surface, such by-products can be extracted by means of a suitable washing procedure. An alternative agent for promoting coupling reactions is 1,10 -carbonyldiimidazole (CDI) (Fig. 2.9) (Bamford et al., 1987). CDI works in a similar fashion to the carbodiimides, activating the carboxylic acid group, thereby increasing its susceptibility towards nucleophilic attack. CDI can be used to form esters between carboxylic acids and alcohols (e.g. PEG) and proceeds via the route illustrated in Fig. 2.10. As shown, the first step involves the loss of CO2 and one equivalent of imidazole to form an acylimidazole, which upon addition of alcohol is converted to an ester with further loss of imidazole. When CDI is reacted with hydroxyl groups (in the absence of carboxylic acid groups) the intermediate can undergo further reaction with an alcohol to form carbonate linkages (Fig. 2.11) (Bamford et al., 1986). However, increased temperatures, strongly basic conditions or the presence of strong nucleophiles such as sodium ethoxide (NaOEt), are often required to promote efficient reaction (Staab, 1962). We have utilised CDI to functionalise the surface of commercially available soft contact lenses with PEG (in house data). Initially, mPEG was reacted with CDI in dichloromethane (CH2Cl2) at room temperature for 1 h (Fig. 2.12). This CDI activated mPEG was then incubated with a commercially available soft contact lens in sodium tetraborate buffer at room temperature for 16 h. mPEG was successfully linked to surface hydroxyl groups via carbonate bonds.

2.10 Ester formation promoted by CDI.

2.11 Carbonate formation promoted by CDI.

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2.12 CDI activation of mPEG.

An alternative compound for promoting carbonate linkages (as well as amides, esters and carbamates) is N,N0 -disuccinimidyl carbonate (DSC), which can be employed under relatively more mild conditions and affords moderate to good yields (Fig. 2.13) (Alsina et al., 1998). A multi-step approach for binding PEG onto the surface of soft contact lenses, with the aim of immobilising ophthalmic drugs, has been proposed by Danion et al. (2007). This process utilises DSC and EDC/NHS and again makes use of existing hydroxyl groups, present on the surface of the contact lens, to direct covalent binding. In the first step, PEI was linked onto surface hydroxyl groups using DSC as the coupling agent to provide surface amine groups (Fig. 2.14). In the second step, NHS activated PEG which contained Biotin at one terminus, was grafted onto the PEI layer via amide bonds using EDC and NHS chemistry. Although multi-step, this approach provides a suitable means of covalently modifying the surface of contact lenses, making use of existing surface functionalities. Clearly, a range of coupling agents are available for directing covalent binding of PEG onto surfaces and offer the possibility of functionalising the surface, PEG or both in order to achieve a more efficient reaction. Another approach for achieving proficient coupling is to convert the carboxylic acid group into a reactive derivative, such as an acid chloride or an acid anhydride.

2.13 N,N0 -disuccinimidyl carbonate (DSC).

2.14 DSC coupling of PEI to surface hydroxyl groups.

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This again serves to increase the electrophilicty of the carboxylic acid group, making it more susceptible to nucleophilic attack.

2.2.5

Carboxylic acid derivatives

Kuang et al. (2008) have used this approach to functionalise the surface of lanthanum hydroxide nanowires. Initially, the surface was made more reactive by introducing surface amine groups. This was achieved through the use of a silane coupling agent, namely (3-amino-propyl)trimethoxysilane (APTMS). mPEG was then converted to an acid chloride by first reacting with succinic anhydride to provide a terminal carboxyl group. This was then converted to an acid chloride through reaction with excess thionyl chloride (SOCl2) at 70 ëC for 5 h (Fig. 2.15). An acylation reaction between surface amine groups and the mPEG-acid chloride was then carried out in the presence of NEt3 and DMF to afford an amide bond (Fig. 2.16). A dramatic increase in water solubility was observed as a result of mPEG grafting onto lanthanum hydroxide nanowires. In order to improve the surface properties of poly(acrylonitrile-co-maleic acid) (PANCMA) asymmetric membranes, Xu et al. (2005) have also made use of a carboxylic acid derivative to covalently link PEG to the surface. In this case, membranes were first refluxed in acetic acid for 1 h to convert surface carboxylic acid groups into acid anhydride groups. The membranes were then immersed in PEG at 100 ëC under nitrogen for 16 h. Washing with excess deionised water removed any unreacted PEG and the membranes were then dried at 40 ëC in a vacuum oven to afford the final grafted assemblies. In this procedure, PEG was covalently linked to the surface via an ester bond.

2.15 Formation of mPEG-acid chloride derivative.

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2.16 Covalent binding of mPEG via an acylation reaction.

As highlighted, carboxylic acid derivatives can provide a suitable means for covalently attaching PEGs onto a surface. In the first case, silanisation was used to functionalise the surface, thereby providing surface amine groups that could react with the acid chloride-mPEG derivative. An alternative approach, which is particularly suitable for grafting PEG onto surfaces with oxide layers, is to employ silanated derivatives of PEG. In this way, surface functionalisation is not required and PEG is immobilised via a siloxane tether.

2.2.6

Siloxane tethers

The adsorption of protein on nanoporous alumina membranes has been reduced by covalently grafting on PEG via a PEG-silane couple (Popat et al., 2004). In this procedure, PEG was first reacted with silicon tetrachloride in the presence of NEt3 to yield PEG-OSiCl3. This was then reacted with trace level hydroxyl groups, present on the alumina membrane surface, for 1 h to form a network of Si-O-Si bonds, thereby immobilising PEG on the surface. Jo and Park (2000) have investigated the covalent binding of silanated PEG derivatives onto glass surfaces as a means of preparing a protein repellent interface. First, two trialkoxysilylated PEG derivatives were prepared (Fig. 2.17), which were then immobilised onto glass surfaces. Fibrinogen adsorption was reduced by more than 95% in comparison to untreated controls. Amphiphilic PEG-silanes possessing siloxane tethers of various lengths (Fig. 2.18) have been grafted onto oxidised silicon wafers by Murthy et al. (2009). Oxidised silicon wafers were placed in a sealed jar containing the specific PEG-silane in toluene and shaken for 12 h. After this time, the wafers were removed and annealed in a vacuum oven at 150 ëC for 12 h. Sonication and rinsing with ethanol and deionised water, followed by drying under a stream of nitrogen, afforded the PEG grafted wafers. As a result of increased steric hindrance, a reduction in chain density of PEO-silanes grafted onto the surface was observed when the siloxane tether length was increased.

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2.17 Trialkoxysilylated PEGs investigated by Jo and Park (2000).

2.18 PEG-silanes possessing siloxane tethers of various lengths.

Amongst the various strategies adopted to direct covalent attachment of PEG onto surfaces, epoxides have also been readily employed.

2.2.7

Epoxides

Epoxides are highly strained cyclic ethers, which undergo reaction with nucleophiles such as alcohols and amines to form ether bonds and secondary amines, respectively. Therefore, epoxides can be used to direct covalent attachment of PEG onto surfaces. Emoto et al. (1996) have coupled a homobifunctional diglycidyl-PEG derivative (Fig. 2.19) onto an aminated quartz surface under various conditions to afford PEGylated surfaces. A similar approach has been adopted by BergstroÈm et al. (1995) to covalently link monofunctional PEG glycidyl ethers onto aminated poly(styrene) (PS). PS was first oxidised using KMnO4/H2SO4 to afford carboxyl and aldehyde groups on the surface. PEI was then covalently coupled to these groups to afford surface amine groups. PEG-epoxide, prepared through the reaction of PEG with

2.19 Homobifunctional diglycidyl-PEG ether.

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2.20 Preparation of PEG-epoxide.

epichlorohydrin in the presence of sodium hydroxide (NaOH) (Fig. 2.20), was then reacted with surface amine groups to afford the PEG functionalised surface. Piehler et al. (2000) have utilised an epoxy-moiety to covalently couple homobifunctional PEGs onto silica surfaces. In this case, surface silanisation with glycidyloxipropyltrimethoxysilane (GOPTS) was carried out first to provide surface reactive epoxide groups. Diamino- and dicarboxy-PEGs were then covalently coupled via amine and ester linkages, respectively. An interesting facet of this approach is the fact that no solvent was employed during the immobilisation step. Alternatively, following silanisation, PEG derivatives were melted onto the surface at 75 ëC for 36 h. Holmberg et al. (1997) have utilised epoxide chemistry to investigate the covalent grafting of PEG to solid PE. Two covalent grafting procedures were performed, whereby a nucleophilic PEG derivative was coupled to electrophilic surface groups or an electrophilic PEG derivative was bound to nucleophilic surface groups. In the first procedure (Fig. 2.21), nucleophilic mono(3-amino-2-hydroxypropyl)-monomethyl ether (amino-PEG) was coupled to the electrophilic surface (prepared by photografting glycidyl methacrylate onto PE plates) by immersing freshly prepared epoxy-functional PE plates in aqueous buffer, pH 9, containing amino-PEG (0.1 g per 10 mL solution) for 2 h at room temperature. The plates were then rinsed extensively with water and dried. In the second procedure (Fig. 2.22), electrophilic PEG diglycidyl ether (PEGepoxide) was coupled to a nucleophilic amino-functional PE surface in aqueous

2.21 Nucleophilic PEG coupled to electrophilic surface.

2.22 Electrophilic PEG coupled to nucleophilic surface.

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buffer, pH 9, under stirring for 2 h. The plates were then rinsed and exposed to aqueous 1% perchloric acid for 30 min in order to open remaining oxirane rings. The plates were then extensively rinsed with water and dried. In terms of grafting density and efficiency, the two approaches were shown to be comparable. An alternative approach to forming ether linkages between PEG and surfaces has been investigated by Litauszki et al. (1996). PEG alkoxide salts were used, whereby mPEG was converted to sodium poly(oxy-ethylene) glycolate (NaPEG) by reaction in the melt with solid sodium or sodium methoxide (NaOMe). The PEG alkoxide was then linked via ether bonds onto halogenated surfaces (Williamson ether synthesis). This approach provides a simple means of attaching PEG to surfaces without the need for linking functionalities or coupling compounds. However, a major drawback of this method is that it is only effective for surfaces that are halogenated.

2.3

Technologies and applications

As demonstrated, a multitude of different covalent binding strategies exist for modifying the surfaces of biomaterials. Clearly, these techniques may be applied to various functional moieties, not only PEG and may be used in conjunction with other surface modification techniques to afford more desirable surface properties. Many studies have been carried out on covalent binding, using a wide range of surfaces and attachment molecules. Indeed, numerous examples are readily available in academic literature; however, a limited number have found successful use in commercial applications. One area, in which covalent binding has proved useful, is in preventing calcific degeneration of prosthetic devices. The calcification of glutaraldehydetreated bioprosthetic heart valves represents a significant problem, which can lead to the clinical failure of such devices. The covalent coupling of an anticalcification agent, 2-amino oleic acid (AOA) (Fig. 2.23), onto the surface of such prostheses, has been shown to inhibit aortic valve calcification (Chen et al., 1994). The covalent binding strategy makes use of free aldehyde groups, which are present on the surface of the heart valve (as a result of glutaraldehyde pretreatment), to react with the secondary amine group of AOA. Covalent coupling occurs via Schiff base formation (Girardot, 1990). This technology is currently

2.23 2-amino oleic acid (AOA).

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employed on Medtronic's FreestyleÕ Aortic Root Bioprosthesis, a stentless valve designed to emulate the native aortic valve. The stabilising properties of covalent binding are also being applied to the field of tissue closure, for the prevention of blood and fluid leakage following surgery (Fortune et al., 2009). The TissuePatchDuralTM surgical film (Tissuemed Ltd) is based on a multi-layered system, constructed from synthetic absorbable polymers. The surface of the film displays functionalised carboxylic acid groups, which promote covalent binding (via amide bond formation) between the film and the tissue surface, thereby sealing the wound. Clinical applications for this technology include sealing air leaks in lung surgery, preventing cerebrospinal fluid leaks and sealing wounds on a variety of organs during general surgery. One advantage of using covalent binding is that it allows biofunctional molecules to be immobilised onto the surface of biomaterials. An example of where this approach has been adopted is on the GenousTM R stent (OrbusNeich), which utilises covalent binding to immobilise murine monoclonal anti-human CD34 antibodies onto the surface of stainless steel stents (Aoki et al., 2005). These antibodies target endothelial progenitor cells, which circulate throughout the entire vascular system and are inherently programmed to mature into endothelial cells. This technology aims to reduce the adverse problems associated with coronary stenting by facilitating the natural healing response and promoting complete endothelialisation of the stent surface.

2.4

Conclusions and future trends

The strategy of covalent binding offers several advantages over alternative techniques, including flexibility, reproducibility and stability. Clearly, the approach and functional moiety adopted is dependent on the specific application in question; however, the diversity offered by covalent binding means that this approach is a particularly attractive option for modifying the surface of biomaterials. Undoubtedly, covalent binding will continue to be used to achieve this goal and this technique therefore has the potential to impact upon a wide range of areas, including cardiology, ophthalmology and diagnostics. It is therefore crucial, that novel immobilisation strategies continue to be investigated, in order to overcome any possible shortcomings that may be associated with current methodologies. Sun et al. (2008) have recently exploited an aqueous based Diels-Alder reaction, to immobilise biomolecules onto N-malemide functionalised glass substrates. A heterobifunctional PEG, containing a cyclopentadiene group at one end and a bioactive molecule at the other (Fig. 2.24), was utilised as the key conjugate.

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2.24 Biomolecule-PEG-cyclopentadiene conjugate.

2.25 Immobilisation of biomolecules through an aqueous based Diels-Alder reaction.

The immobilisation step (Fig. 2.25) was achieved by incubating this conjugate with the functionalised glass substrate in water at room temperature for 12 h, followed by washing with deionised water. A diverse range of biomolecules were immobilised using this technique, which benefits from mild reaction conditions, high efficiency and selectivity. This technique may therefore prove useful for functionalising the surfaces of diagnostic and therapeutic medical devices, such as sensors. Another promising area for directing covalent binding is that of `click chemistry'. Kolb et al. (2001) defined `click chemistry' as a modular, high yielding reaction that can be carried out in a benign solvent (such as water). This approach therefore offers wide scope for the development of highly efficient reaction methodologies, which may prove useful for the modification of biomaterial surfaces. Based on this approach, Fournier et al. (2009) have developed an efficient strategy for attaching PEG onto the surface of PU films and foams, which makes use of a copper catalysed Huisgen 1,3-dipolar cycloaddition. The azide functionalised mPEG (Fig. 2.26) was combined with alkyne-functionalised PU in the presence of water and a copper catalyst. The reaction was performed overnight at room temperature.

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2.26 Azide-functionalised mPEG used in Huisgen 1,3-dipolar cycloaddition.

The successful immobilisation of mPEG was confirmed by an increase in surface wettability when compared with an untreated control. This mild and highly selective approach may therefore offer advantages over conventional methods for the surface modification of PU, as well as other materials.

2.5

References

Ademovic Z, Klee D, Kingshott P, Kaufmann R and HoÈcker H (2002), `Minimisation of protein adsorption on poly(vinylidene fluoride)', Biomol Eng, 19, 177±82. Alsina J, Rabanal F, Chiva C, Giralt E and Albericio F (1998), `Active carbonate resins: application to the solid-phase synthesis of alcohol, carbamate and cyclic peptides', Tetrahedron, 54, 10125±52. Aoki J, Serruys P W, van Beusekom H, Ong A T L, McFadden E P, Sianos G, van der Giessen W J, Regar E, de Feyer P J, Davis H R, Rowland S and Kutryk M J B (2005), `Endothelial progenitor cell capture by stents coated with antibody against CD34', J Am Coll Cardiol, 45, 1574±9. Bailey F E and Koleske J Y (1976), Poly(ethylene oxide), New York, Academic Press. Bamford C H, Middleton I P and Al-Lamee K G (1986), `Studies of the esterification of dextran: routes to bioactive polymers and graft copolymers', Polymer, 27, 1981±5. Bamford C, Middleton I, Al-Lamee K, Paprotny J and Satake Y (1987), `Routes to bioactive hydrophilic polymers', Polym J, 19, 475±83. È sterberg E, Holmberg K, Hoffman A S, Schuman T P, Kozlowski A and BergstroÈm K, O Harris J M (1995), `Effects of branching and molecular weight on surface-bound poly(ethylene oxide) on protein rejection', in Cooper S L, Bamford C H and Tsuruta T, Polymer biomaterials in solution, as interfaces and as solids, Utrect, VSP, 195±204. Byun Y, Jacobs H A and Kim S W (1995), `Heparin surface immobilisation through hydrophilic spacers: thrombin and antithrombin III binding kinetics', in Cooper S L, Bamford C H and Tsuruta T, Polymer biomaterials in solution, as interfaces and as solids, Utrect, VSP, 568±79. Chen H, Chen Y, Sheardown H and Brook M A (2005), `Immobilisation of heparin on a silicone surface through a heterobifunctional PEG spacer', Biomaterials, 26, 7418±24. Chen W, Schoen F J and Levy R J (1994), `Mechanism of efficacy of 2-amino oleic acid for inhibition of calcification of glutaraldehyde-pretreated porcine bioprosthetic heart valves', Circulation, 90, 323±9. Danion A, Brochu H, Martin Y and Vermette P (2007), `Fabrication and characterisation of contact lenses bearing surface-immobilised layers of intact liposomes', J Biomed Mater Res, 82A, 41±51. Emoto K, Harris J M and Van Alstine J M (1996), `Grafting poly(ethylene glycol) epoxide to amino-derivatised quartz: effect of temperature and pH on grafting density', Anal Chem, 68, 3751±7. Fortune D H, Kettlewell G, Mandley D J, Morris D and Thompson I (2009), `Tissueadhesive formulations', US Patent Application 20090018575. Fournier D, De Geest B G and Du Prez F E (2009), `On-demand click functionalisation of polyurethane films and foams', Polymer, 50, 5362±7.

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Girardot J-M (1990), `Prevention of prosthesis calcification', US Patent 4976733. Goddard J M and Hotchkiss J H (2007), `Polymer surface modification for the attachment of bioactive compounds', Prog Polym Sci, 32, 698±725. Graham N B (1987), `Poly(ethylene oxide) and related hydrogels', in Peppas N A, Hydrogels in medicine and pharmacy, Vol. 2. Polymers, Boca Raton, FL, CRC Press, 95±113. Han D K, Park K D, Ahn K-D, Jeong S Y and Kim Y H (1989), `Preparation and surface characterisation of PEO-grafted and heparin-immobilised polyurethanes', J Biomed Mater Res, 23, 87±104. Holmberg K, Tiberg F, Malmsten M and Brink C (1997), `Grafting with hydrophilic polymer chains to prepare protein-resistant surfaces', Coll Surf A: Physicochem Eng Asp, 123±124, 297±306. Janissen R, Oberbarnscheidt L and Oesterhelt F (2009), `Optimised straight forward procedure for covalent surface immobilisation of different biomolecules for single molecule applications', Coll Surf B: Biointerf, 71, 200±7. Jo S and Park K (2000), `Surface modification using silanated poly(ethylene glycol)s', Biomaterials, 21, 605±16. Kingshott P, Thissen H and Griesser H J (2002), `Effects of cloud-point grafting, chain length and density of PEG layers on competitive adsorption of ocular proteins', Biomaterials, 23, 2043±56. Kolb H C, Finn M G and Sharpless K B (2001), `Click chemistry: diverse chemical function from a few good reactions', Angew Chem Int Ed, 40, 2004±21. Kuang J, Yuan J, Zhou M, Yuan W, Sui X and Li Z (2008), `Grafting of PEG onto lanthanum hydroxide nanowires', Mat Lett, 62, 4078±80. Lee J H, Lee H B and Andrade J D (1995), `Blood compatibility of polyethylene oxide surfaces', Prog Polym Sci, 20, 1043±79. Litauszki L, Howard L, Salvati L and Tarcha P J (1996), `Grafting of PEO via the Williamson ether synthesis onto polymeric surfaces and their affinity for proteins', in Ratner B D and Castner D G, Surface modification of polymeric biomaterials, New York, Plenum Press, 183±91. Murthy R, Shell C E and Grunlan M A (2009), `The influence of poly(ethylene oxide) grafting via siloxane tethers on protein adsorption', Biomaterials, 30, 2433±9. Piehler J, Brecht A, Geckeler K E and Gauglitz G (1996), `Surface modification for direct immunoprobes', Biosen Bioel, 11, 579±90. Piehler J, Brecht A, Valiokas R, Liedberg B and Gauglitz G (2000), `A high-density poly(ethylene glycol) polymer brush for immobilisation on glass-type surfaces', Biosen Bioel, 15, 473±81. Popat K C, Mor G, Grimes C and Desai T A (2004), `Poly(ethylene glycol) grafted nonporous alumina membranes', J Membrane Sci, 243, 97±106. Staab H A (1962), `New methods of preparative organic chemistry IV: syntheses using heterocyclic amides (azolides)', Angew Chem Internat Edit, 1, 351±67. Sun X-L, Yang L-C and Chaikof E L (2008), `Chemoselective immobilisation of biomolecules through Diels-Alder and PEG chemistry', Tet Lett, 49, 2510±13. Tseng Y-C, McPherson T, Yuan C S and Park K (1995), `Grafting of ethylene glycolbutadiene block copolymers onto dimethyl-dichlorosilane-coated glass by irradiation', Biomaterials, 16, 963±72. Xu Z-K, Nie F-Q, Qu C, Wan L-S, Wu J and Yao K (2005), `Tethering poly(ethylene glycol)s to improve the surface biocompatibility of poly(acrylonitrile-co-maleic acid) asymmetric membranes', Biomaterials, 26, 589±98.

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Surface modification of biomaterials by heparinisation to improve blood compatibility X . Z H A O and J . M . C O U R T N E Y , University of Strathclyde, UK

Abstract: This chapter presents a brief review on the importance of surface modification of blood contacting materials for improved blood compatibility. Blood compatibility is mainly influenced by the surface properties of the materials. The interaction of the surface and blood leads to a blood coagulation which is not desirable for some medical applications. In order to create a blood compatible material with an anti-thromobogenic or nonthromobogenic surface, many theoretical hypotheses have been postulated. In practice, surface modification by utilisation of heparin is one of the mostly widely accepted approaches for improvement of blood compatibility. In this chapter, some of the immobilisation methodologies are reviewed and the effects of heparinisation on blood compatibility both in vitro and in vivo are also assessed including protein adsorption, platelet adhesion, thrombus formation and other factors which influence blood compatibility. Key words: heparinisation, surface modification, blood compatibility, heparin, blood and surface interaction.

3.1

Introduction

Biomaterials have been the base for the medical device industry. The medical device market is about 50 per cent of the world pharmaceutical market in terms of relative size, but is also growing faster than its drug counterpart. The biomaterial industry has experienced phenomenal growth due to the health care industry and market demands for advanced medical devices. In 2010 the UK market for medical devices is now worth about $8.4 billion [1]. According to a new market research report, `Global biomaterials Market (2009±2014)', published by MarketsandMarkets, the total global biomaterials market is expected to be worth US$58.1 billion by 2014, growing at a CAGR of 15.0% from 2009 to 2014. The US market is expected to account for nearly 42% of the total revenues. The biomaterials market today has already exceeded $28 billion [2]. Biomaterials need to be blood compatible while contacting blood. Blood compatibility is the key consideration factor when biomaterials are designed for blood contacting applications. Blood tends to coagulate when it contacts a

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foreign surface and the surface properties are critical for maintaining the nonthrombogenicity of the surface. The hypothetical correlation of blood compatibility with surface properties can be traced back to 1863, when Lister (1863) experimented with a jugular vein bypass in sheep, comparing rubber and glass tubing. He found that blood clotted more slowly in rubber than in glass [3]. Freund (1885) found that glassware coated with petroleum jelly, such as Vaseline, delayed the blood clotting time [4]. Bordet and Gengou (1903) discovered that the blood clotting time was increased when glass was covered with paraffin wax [5]. These above observations and experiments led to the conclusion that the nature of the polymer surface affects the clotting process. Since the early studies, numerous hypotheses attempting to correlate polymer surfaces with blood compatibility have been proposed, as shown in Table 3.1. From Table 3.1, it is clear that various basic concepts and hypotheses for the development of blood compatible surfaces have been described. It is not surprising to find different views and contradictory results in the literature due to the multivariable nature of the blood compatibility of biomaterials. It is also difficult to regulate blood-foreign surface interaction by any simple hypothesis. Surface modification of biomaterials for improved blood compatibility can be achieved by many means according to variable hypotheses in the literature [6]. However, one of the most widely used clinical practices is the utilisation of heparinised biomaterials for blood-contacting applications.

Table 3.1 Hypotheses for the co-relationship between surface of the material and the blood compatibility Proposers

Hypothesis

Reference

Neubauer and Lampert (1930) Saywer et al. (1953)

Lampert Rule of blood clotting time

[7]

Negatively charged surfaces tend to be nonthrombogenic Antithrombotic agents coating surface Minimal critical surface tension Surface morphology

[8]

Biomembrane-mimetic hypothesis

[13] [14]

Fibrinogen retention hypothesis

Described in [6]

Gott et al. (1963) Zisman (1964) Lyman (1975) and Hecher and Edwards (1981) Hayward and Chapman (1984) and Ishihara et al. (1990) Okkema et al. (1991)

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3.2 Bioactive molecule: heparin Standard heparin is made up of long-chain glycosaminoglycans which have a diverse molecular weight ranging between 5000 and 30 000 [15], consisting of a variably sulphated repeating disaccharide unit as shown in Fig. 3.1. Unique features of heparin are: the presence of N-sulphate groups; the -D anomeric configuration of the glucosaminidic bond; the 1-4 glycosidic bonds. Heparin has long been known to inhibit blood clotting [16]. This anticoagulation effect clinically is via a mechanism, which primarily influences antithrombin III to block the action of thrombin and factor Xa [17]. One unit of bioactivity of heparin (the `Howell unit') is an amount approximately equivalent to 0.002 mg of pure heparin, which is the quantity required to keep 1ml of cat's blood fluid for 24 hours at 0 ëC. Other than standard high molecular weight heparin, there are heparins with low molecular weight and molecules with similar heparin structures, termed heparinoids. Low molecular weight (LMW) heparin refers to a heparin derivative that has been obtained by fractionation or depolymerisation of commercial grade heparin and has a lower average molecular weight than the parent heparin. Certain preparations of LMW heparins, averaging at 4000±6000, are at least as anti-thrombotic as standard heparin but produce less bleeding in animals. Heparinoids refer to molecules having heparin-like structures and properties, such as chondroitin sulphates, dermatan sulphate and heparitin sulphate. Physiological functions of heparin, other than the clinical use as an anticoagulant, have been continuously discovered. It was found that heparin is able to inhibit the proliferation of vascular smooth muscle cells [18] and also has the function to suppress delayed hypersensitivity [19]. The discovery of a new function of heparin, independent of anticoagulation, is the involvement of heparin in the regulation of angiogenesis, which may eventually lead the new use of heparin in new blood vessel formation for wound healing [20]. Damus et al. (1973) [21] suggested that heparin-like molecules present on the vascular endothelium endowed these surfaces with thromboresistant properties. These molecules are synthesised by vascular tissues. Therefore, to mimic the natural blood vessel with a nonthrombogenic surface, the modification of surface of biomaterials using heparin for improved blood compatibility is one of the widely accepted clinical approaches.

3.1 Chemical structure of heparin.

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3.3

59

Blood±biomaterial interaction

When blood contacts a biomaterial surface in medical applications, such as haemodialysis, haemoperfusion, haemofiltration and cardiopulmonary bypass, the following groups of events occur: 1. Adsorption of plasma proteins onto the polymer surface. 2. Activation of the complement and kinin/kallikrein systems, blood cells and intrinsic coagulation, initiated by the adsorbed proteins from the system. 3. Adhesion of cell components (thrombocytes, granulocytes and monocytes) to the protein coating. 4. Formation of fibrin onto the surface and also possible activation of the fibrinolytic system.

3.3.1

Protein adsorption

It is commonly stated that the first observable event at the interface between a foreign material surface and blood is adsorption of blood proteins onto the material. This occurs within a few seconds of blood coming into contact with the material. The proteins adsorbed initially are those present in relatively high concentration in the plasma, including species such as albumin, fibrinogen, globulin and IgG. However, these adsorbed proteins will eventually be replaced by trace proteins, such as the coagulation proteins of the intrinsic and extrinsic pathways, complement proteins and fibrinolytic proteins [22], by high molecular weight kininogen (HMWK) and Factor XII (Hageman factor). This is called the `Vroman effect' [23]. It is also believed that the adsorption rate, quantitative composition of the adsorbed protein layer, protein conformation changes in the spatial architecture, and protein globular structures on the polymer surface define the subsequent events [24]. It is considered that the information on fibrinogen and albumin adsorption is relevant for blood-biomaterial interactions [25]. There has been considerable work attempting to correlate the blood compatibility of biomaterials with the adsorption of fibrinogen, albumin and other proteins. The adsorbed fibrinogen has been utilised in monitoring the blood response to PVC tubing and was regarded as an index for blood compatibility evaluation [26].

3.3.2

Platelet reactions

Platelet deposition on artificial surfaces is a characteristic feature of platelet function in vitro and in vivo [27]. The attachment and activation of platelets are associated with the adsorption of adhesive proteins, which are capable of binding to the platelet glycoprotein IIb-IIIa (GPIIb-IIIa) complex and other receptors of stimulated platelets. This complex is able to recognise specifically the Arg-Gly-Asp(RGD)-tripeptide sequences of adhesive proteins, such as

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fibrinogen (FGN), fibronectin (FN), vitronectin (VN) and von Willebrand factor (vWF). After deposition, platelets undergo morphological changes and contractions, and the platelet constituents, such as adenosine diphosphate (ADP), serotonin and other substances, are excreted. These released constituents influence further platelet adhesion and aggregation on the surface.

3.3.3

Coagulation

Thrombus formation occurs either intrinsically by surface-mediated reactivity or extrinsically through factors derived from tissues. The two systems converge upon a final common path, which leads to the formation of an insoluble fibrin gel when thrombin acts on fibrinogen. For blood-biomaterial interactions, the intrinsic coagulation system is the more important [27]. From the biochemistry point of view, there are at least 12 plasma proteins, which interact in a series of reactions leading to blood clotting. Their biochemical properties are summarised in Table 3.2 and the clotting mechanism has been well documented [27].

Table 3.2 Human proteins of coagulation system Protein (clotting factor)

Intrinsic system Factor XII Prekallikrein High MW kininogen Factor XI Factor IX Factor VIII Vwf Extrinsic system Factor VII Tissue factor Common pathway Factor X Factor V Prothrombin Fibrinogen Factor XIII

Molecular weight (no. of chains)

Normal plasma concentration (g/ml)

80,000 (1) 85,000 (1) 105,000 (1) 160,000 (2) 68,000 (1) 265,000 (1) 1±15,000,000

30 50 70 4 6 0.1 7

47,000 (1) 46,000 (1)

0.5

56,000 (2) 330,000 (1) 72,000 (1) 340,000 (6) 320,000 (4)

10 7 100 2500 15

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Active form

Serine protease Serine protease Cofactor Serine protease Serine protease Cofactor Cofactor for platelet adhesion Serine protease Cofactor Serine protease Cofactor Serine protease Clot structure Transglutaminase

Heparinisation to improve blood compatibility

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Fibrinolysis

The fibrinolytic system removes unwanted fibrin deposits and facilitates the healing process after injury and inflammation. It is a multicomponent system composed of precursors, activators, cofactors and inhibitors, and has been studied extensively [25].

3.3.5

Erythrocytes and leucocytes

Erythrocytes (red cells) make up 96% of the total blood cell volume but, in discussing blood cell-foreign material surface interactions, relatively little attention has been paid to erythrocytes. It is known that this interaction may result in the release of several lipids and fatty acids from the red cell membrane, causing erythrocyte adhesion, a significant change in cell (membrane) metabolism, or haemolysis [28]. The adhesion of red cells to a foreign material surface is mainly determined by the nature of the surface, e.g. chemical structure and roughness, and blood flow parameters, e.g. shear rate, along the surface [29]. The reaction of leucocytes, (white blood cells), is directly related to the host defence of the human body towards foreign material surfaces. The interaction of white blood cells and foreign material surfaces is mediated by the adsorbed protein layer at the surface and depends also on the type of white blood cell. In blood-contacting applications, the leucocyte response is often linked to the relationship between leucocytes and complement activation [27].

3.3.6

Complement activation

Activation of the complement system caused by blood-material interaction occurs through the alternative pathway [30]. The complement system consists of about 20 proteins, which can be stimulated by bacteria and endotoxins, but also by a foreign material surface, such as a dialysis membrane or blood tubing. It is part of the immune system response to a foreign material. Complement activation is considered an important parameter for characterising the blood compatibility of biomaterials, particularly extracorporeal applications. For dialysis membranes, the decision to develop and use new membranes mainly follows their reduced capacity to induce complement activity in clinical dialysis. Triggering of complement activation only depends on the chemical composition of the foreign material surfaces, which is relevant to the affinity to C3b. In principle, a hydrophilic hydroxyl group gives rise to low coagulation and platelet adhesion, but stimulates markedly complement activation. Cationic surfaces trigger platelet adhesion, but demonstrate a low influence on the activation of complement and coagulation [31].

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3.3.7

Surface modification of biomaterials

Interrelationships

As expected, blood-biomaterial interactions are very complicated, and there are many interrelated reactions and feedback networks [32]. For example, platelet reactions are interrelated with the coagulation system to promote thrombin formation and can interact with the fibrinolytic system by the binding of plasminogen to the GPIIb-IIIa complex [33]. It is found that leucocytes are involved in the intrinsic coagulation, fibrinolysis and complement activation. The leucocyte membrane contains phospholipids, which may play a role in blood clotting via the intrinsic pathway [34]. To assess the blood compatibility of a biomaterial, many factors need to be considered including: the patient, the blood, the bulk and surface properties of the material, the blood flow condition, and the amount of heparin in the blood. The challenge facing to both the research community and industry is to design a new biomaterial with perfect blood compatibility. Surface modification of an existing biomaterial, with an acceptable performance, in order to achieve improved blood compatibility for clinical applications, is a practical approach in the medical device industry. This can help to produce the compatible materials for devices such as catheters, biosensors, heart valves, vascular prostheses, dialysis equipment and other devices associated with blood circulation systems.

3.4

Surface modification by heparinisation for improved blood compatibility

3.4.1

Definition

Surface modification by heparinisation is a process for modifying a polymeric biomaterial with heparin to obtain improved blood compatibility. The process is also termed heparin immobilisation and the modified surface is called a heparinised surface. Based on the chemical nature of heparin, the molecular design of a heparinised surface can be achieved.

3.4.2

Methodological approaches to carry out the heparinisation process

The heparin molecule contains hydroxyl (±OH), carboxyl (±COOH), amino (±NH2), and sulphate (±SO3ÿ). The reactivity of heparin is based on the reaction and interaction between these functional groups and functional groups at the surface of material. Immobilisation of heparin on a polymeric surface can be generally described according to Fig. 3.2. A typical heparin surface immobilisation procedure is divided into two steps:

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3.2 Immobilisation of heparin onto a polymer surface.

Step 1: Study of the chemical structure of the substrate surface, cleaning the surface; modification of the surface, if necessary, to introduce reactive groups. Step 2: React between heparin and the surface (or modified) and purify to obtain heparinised surface. In literature, the surface heparinisation can be achieved by: · ionic binding of heparin · covalent binding of heparin · physical blending of heparin for control release. Ionic binding of heparin As heparin is of strong anionic character, it can bind to cationic surfaces to form ionic complexes. In 1963 Gott et al. first reported a process for coating a

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graphite layer onto a solid polymer surface, followed by adsorption of a quaternary ammonium salt, benzalkonium chloride, on to the graphite layer. After the ionic complexation, a heparin modified surface, designated graphitebenzalkonium-heparin (GBH), was obtained [9]. The most widely used method for ionic binding heparin involves a pretreatment of a polymer surface with a chemical solution, such as tridodecylmethylammonium chloride (TDMAC) in a solvent at 1±5% concentration. After the pretreatment, heparin can be bound to the surface [35]. Merker et al. quaternised silica filler for silicone rubber, using -aminopropyltriethoxysilane (APTES) followed by treatment with heparin [36]. However, these surface binding technologies cannot create a stable heparinised surface. Treatment of cellulose dialysis membranes such as membrane, with ethylene imine and subsequent treatment by heparin produced anti-thrombogenic materials [37]. In order to create a stable heparinised surface, heparin was firstly bound to the polymer surface by ionic complexing, followed by crosslinking using a crosslinking agent such as glutaraldehyde [38]. Covalent binding of heparin In order to create a stable heparinised surface, heparin can be covalently bound to a polymer surface. This normally involves an introduction of functional groups at the surface. For example, heparin can be covalently bound to polyurethane catheters via three steps [39]: Step 1: Hydrolysis of polyurethane in 3 M NaOH in the presence of isopropanol to introduce hydroxyl and amine groups. Step 2: React free amine groups with phosgene and diamino alkane spacers were coupled to the isocyanate groups. Step 3: The terminal amine groups of the immobilised spacer were used for coupling with 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) activated heparin and a heparinised surface thus was formed at the level of 2±4 g/cm2. Heparin can also be covalently bound to collagen using EDC according to the following steps [40]: Step 1: Heparin and EDC are dissolved in water. Step 2: Collagen is added in the solid form, i.e. film, powder or fibre. Step 3: Change the pH to 4.8 to start the immobilisation process, further purify the collagen with water to obtain a heparinised collagen surface. Using a PEO spacer to immobilise heparin can enhance the bioactivity of the bound heparin. A typical method is shown as follows [41±43]:

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Step 1: Activation of polymer surfaces to introduce reactive groups. Step 2: Coupling PEO-diamines (Jeffamine) or PEO-OH to the surface to introduce PEO spacer with reactive groups. Step 3: Using EDC as coupling agent to react with heparin or end-point method to bind heparin. The heparin activity was measured on the basis of prolongation of activated partial thromboplastin time (APTT). The activity of immobilised heparin was observed after the amino groups at the terminal end of the spacer molecule, not used for immobilization of heparin, were blocked with formaldehyde. The flexible PEO spacer appeared to relieve the steric hindrance by the carrier to the formation of a heparin complex with cofactor AT III. The activity of immobilised heparin increased with elongation of the spacer chain and the highest activity was obtained when the chain length of the spacer was between 10 and 20. Longer spacer groups decreased the activity because the high mobility of the hydrated PEO chain hindered AT III access to the immobilised heparin. A dense albuminated layer, a heparinised layer, and a mixed layer on a poly(acrylic acid)-grafted surface can be achieved via visible light induced photopolymerisation. The procedure comprises three reaction steps [44]: Step 1: By visible light irradiation, acrylic acid (AA) was graft-polymerised on a segmented polyurethane (SPU) film that was preimpregnated with camphorquinone. Step 2: Adsorption of multiply styrenated albumin or styrenated heparin or their mixture, followed by visible light irradiation in the presence of carboxylated camphorquinone. Step 3: Covalent bonding between polyAA graft chain and polymerised biomacromolecule and between polymerised biomacromolecule to ensure the formation of a stable immobilised multilayer. Platelet adhesion was markedly reduced on polymerised albuminated, polymerised heparinised, and copolymerised layers, whereas adhesive and proliferative potentials of endothelial cells, which were comparable to those of commercial tissue culture dishes, were observed on these surfaces. Coimmobilisation of fibronectin and basic fibroblast growth factor enhanced these potentials. These densely multilayered surfaces may be suitable for artificial and tissue-engineered devices. Michanetzis et al. [45] selected four commercially available biomaterials for heparin immobilisation: polydimethylsiloxane (silicone rubber), polyvinylchloride (plasticiser: di-(ethyl-2hexyl)-phthalate (DEHP)), polyethylene (low density) and polypropylene. All materials were provided by Rehau (Germany) in tube format with an internal diameter of 2 mm, and were cut 30 cm long. Two heparinisation techniques, one indirect and one direct, were selected to bind covalently heparin molecules on a biomaterial surface according, to methods

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developed by Bamford and Al-Lamee [46, 47] and Seifert and Groth [48], respectively. Two tubes per material were used for each method. The first heparinisation method was applied in three steps: 1. Hydroxylation of the surface: the polymers were treated with a 10% aqueous solution of potassium peroxydisulphate (used as the oxidising agent) for 1 h at 701 ëC dynamically, using a Harvard Apparatus syringe pump at a shear rate of 200 s1 infusion-withdrawal mode. The polymers were then washed copiously with hot water and dried in vacuum. 2. Grafting to functionalised polymers: after the above treatment the polymers were submitted to grafting reactions in aqueous solution with the aid of the ceric ion technique [49] using acrylamide (10% w/v) in nitric acid and ceric ammonium solution (0.04 and 0.002 m, respectively) at 501 ëC for 1 h under a stream of nitrogen, creating free amino groups on the materials surface. Materials were washed extensively with 0.05 m NaOH followed by hot water. 3. Heparin immobilisation on the treated materials surface using 1% w/v heparin solution (heparin sodium, Serva, 174 000 IU/g) at pH 4.5 containing 0.1% w/v carcodiimide hydrochloride (EDC-Sigma) as the coupling agent. The reaction was continued for 48 h at ambient temperature under static conditions. The second method heparin immobilisation was performed using heparin sodium (Serva, 174 000 IU/g) with glutaraldehyde (Sigma) as the coupling agent, by modification of the method introduced by Peppas and Merill [50]. A 1:1 mixture of 12% aqueous solution of heparin and 6% aqueous solution of glutaraldehyde adjusted with 0.1 m sulphuric acid to pH 5.2 was incubated in the tubes for 2 h under static conditions. Another approach is to introduce carboxyl groups via plasma glow discharge onto polyurethane material and then heparin is covalently bound to the surface, according to Figs 3.3 and Fig. 3.4 [51].

3.3 Schematic diagram showing graft copolymerisation on a PU surface.

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3.4 Schematic diagram showing PEO grafting and heparin immobilisation [with permission 51].

Surface heparinisation with Carmeda BioActive SurfaceTM (CBASTM) Heparin is bound to the surface by End-point attachmentTM [52]. This proprietary coupling technique is the key to non-thrombogenicity and longterm stability. The CBAS coating process is entirely water based. In the initial steps, a layer-by-layer base matrix of a positively charged polymeric amine and a negatively charged sulphated carbohydrate is built up. Finally heparin is covalently bound to this base matrix by End-point attachment, at one terminus of the heparin molecule. The coating thickness varies depending on the specific coating configuration (based on application and substrate). A typical coating thickness is in the range of 100 nm. Zhao et al. [53] studied medical grade plasticised poly (vinyl chloride) (PVC) in tubing form for surface modification by utilisation of the CBAS heparinisation technology, as shown in Fig. 3.5. For understanding the relationship between the artificial surfaces and blood, electron spectroscopy for chemical analysis (ESCA) technology was employed to determine the surface chemical composition and bonding structure. Physical blending of heparin for control release Polyurethane was blended with silk fibroin and heparin to prepare a heparinreleasing system. The release rate and the percentage of the cumulative amount of the released heparin can be controlled by the loading amount of heparin in the film, the composition ratio of silk fibroin to polyurethane, and the thickness of the film. The slower and more sustained release of heparin can be obtained by increasing the film thickness, the loading amount of heparin and the content of

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3.5 Surface modification of PVC-DEHP using end-point attachment of heparin [with permission 53].

silk fibroin in the film. Thus, high bioactivity and long lasting antithrombogenicity of the raw heparin can be maintained for the blended film. Coagulation time tests showed that the composite film had good blood compatibility [54].

3.4.3

Analysis of heparinisation degree

The amount of immobilised heparin was estimated by a spectrophotometric assay similar to that introduced by Farndale [55], described in detail for heparin determination by Van de Lest et al. [56], using dimethyl methylene blue dye (Serva) for colour development and measuring optical density at 600 nm with a TECAN Spectra Classic 96-well plate spectrophotometer (TECAN, Austria). The amount of heparin immobilised on the polymer surfaces was also determined using the toluidine blue method as previously reported [57, 58]. For example, toluidine blue (25 mg, Fluka Chemie, Buchs, Switzerland) was dissolved in 0.01 N hydrochloric acid containing 0.2 wt% NaCl. A known amount of a heparin aqueous solution (2 ml) was added to the toluidine blue solution (3 ml) and the mixed solution was agitated using a vortex mixer. n-Hexane (3 ml) was then added, and the mixture was well shaken so that the toluidine blue-heparin complex was extracted into the organic layer. The unextracted toluidine blue which remained in the aqueous phase was determined by measuring the absorption at 631 nm. The linear relationship between the

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absorbance at 631 nm caused by the residual toluidine blue and the concentration of heparin in the aqueous solution was obtained and used as a calibration curve to determine the amount of immobilised heparin. To ascertain the amount of immobilised heparin, the toluidine blue solution (3 ml) was mixed with an aqueous solution (2 ml) and the heparin immobilised PU polymer (1  1 cm2) was immersed in the mixed solution for 30 min. n-Hexane (3 ml) was then added to ensure a uniformity in treatment and the mixture was shaken well. After the removal of the polymer, the aqueous layers of the solution were sampled. The absorbance at 631 nm was measured and the amount of immobilised heparin was then calculated from the previously constructed calibration curve.

3.4.4

Evaluate the effect of heparinisation on blood compatibility

Effect on thrombus formation A heparin immobilised surface is generally considered to be capable of enhancing the blood compatibility of that surface, reducing platelet adhesion, loss of blood cells, and increasing plasma recalcification time and activated partial thromboplastin time (APTT). The mechanism by which heparin inhibits blood clotting is believed to be associated with the fact that the immobilised heparin, unlike soluble heparin, also inhibits initial contact activation enzymes through an antithrombin-mediated pathway, and thus has enhanced anticoagulant properties [59]. For example, in order to develop collagen for stent coating or fabrication of vascular grafts, collagen was heparinised by crosslinking collagen with extensively periodate oxidised heparin and/or by covalently bonding of mildly periodate oxidised heparin. Both ways of heparinisation have no effect on platelet adhesion and could not abolish induction of platelet procoagulant activity. However, thrombin generation was completely prevented under static and flow conditions. The functionality of immobilised heparin was confirmed by specific uptake of antithrombin III, 13:5  4:7 pmol/cm2 and 1:95  0:21 pmol/ cm2 for mildly and heavily periodated heparin, respectively. It indicates that immobilisation of heparin on collagen, even as a crosslinker, is a very effective way of preventing surface thrombus formation [60]. Heparinisation of artificial surfaces has been proven to reduce the intrinsic thrombogenicity of such surfaces. However, the mechanism by which immobilised heparin reduces thrombogenicity is not completely understood. Keuren et al. [61] studied heparin-, alginic acid- and chondroitin-6-sulphatecoated surfaces for protein adsorption, platelet adhesion and thrombin generation. The protein-binding capacity from solutions of purified proteins was significantly higher for heparin-coated surfaces when compared with alginic

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acid- and chondroitin sulphate-coated surfaces. Yet, when the surfaces were exposed to flowing plasma, only the heparinised surface adsorbed significant amounts of antithrombin III. None of the surfaces adsorbed fibrinogen under these conditions, and as a result no platelets adhered from flowing whole blood. These results indicate the potential problems when protein adsorption and platelet adhesion from anticoagulated blood are used to assess the thrombogenicity of (coated) artificial surfaces. Indeed, the thrombin generation potential of the different surfaces varied remarkably: while non-coated surfaces readily produced thrombin, alginic acid- and chondroitin sulphate-coated surfaces showed a marked reduction and virtually no thrombin was generated in flowing whole blood passing over heparinised surfaces. In an effort to better mimic the thromboresistive nature of the vascular endothelium, extracorporeal circuits bonded with heparin or phospholipids were developed and tested in vivo [62]. Using no systemic heparinisation, these circuits were compared with standard poly(vinyl)chloride (PVC) (Tygon) in a rabbit model of extracorporeal circulation (ECC). Control circuits were run, with and without systemic heparinisation, and used as comparison groups against the test circuits. Two New Zealand White rabbits were used per study: One was used as the platelet donor for 111 indium platelet labelling; the other animal was placed on bicaval ECC for 4 hours. Circuits (heparin coated n ˆ 6, phospholipid coated n ˆ 8, nonheparinised controls n ˆ 14, heparinised controls n ˆ 18) consisted of 1 m of tubing, two downsizing connectors, and two venous cannulae. ECC blood flow was at least 75 ml/min. Platelet and fibrinogen measurements were made hourly, and circuit dosimetry was performed at the end of the study or on circuit thrombosis. Thrombosis of the circuit occurred in one heparin coated, two phospholipid coated, and eight nonheparinised control circuits. None of the heparinised control circuits thrombosed. There was no significant difference between the groups with regard to platelet count or platelet adhesion. Test circuits exhibited preservation of fibrinogen levels. In this rabbit model of ECC, circuits coated with heparin or phospholipids appeared to preserve fibrinogen levels but did not reduce platelet adhesion or consumption. Platelet adhesion on the cardiopulmonary bypass oxygenator membrane is associated with impaired haemostasis. The effects of heparin coating of the oxygenator membrane on protein adsorption and platelet adhesion on the surface have been assessed, according to the following method: uncoated and heparincoated polypropylene membranes were incubated in whole blood with small(1 U/mL) or large-dose (5 U/mL) heparin as an anticoagulant for 3 h at 37 ëC. The amount of platelets adhering on each fibre was assessed by enzyme immunoassays, using monoclonal antibodies directed against CD42b (GP Ib) and CD61 (GP IIb/IIIa). Platelet activation was assessed by measuring plasma guanosine monophosphate 140 levels. The amount and composition of the adsorbed proteins on the surface were analysed by a bicinchoninic acid protein assay and by sodium dodecyl sulphate-polyacrylamide gel electrophoresis and

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Western blotting technique. Heparin coating of the fibres significantly reduced platelet adhesion on the surface. However, platelet activation was reduced by heparin coating only with small-dose heparinisation. The adsorption of platelet adhesive proteins, such as fibrinogen and von Willebrand factor, was not altered, whereas that of fibronectin was increased by heparin coating. This indicates that heparin coating of the oxygenator fibres can decrease platelet adhesion without affecting adsorption of major adhesive proteins. Surface heparin coating is associated with an increased fibronectin adsorption on the fibres. It can reduce platelet adhesion and activation in the presence of small-dose heparinisation, potentially reducing the inflammatory response and activation of thrombosis and fibrinolysis [63]. Postoperative morbidity after cardiopulmonary bypass most commonly manifests as bleeding diatheses or pulmonary dysfunction. The pathophysiology has been attributed to the activation of cellular and humoral components of blood after contact with an artificial surface. Development of a surface that would be nonthrombogenic and also would constitute a less potent inflammatory stimulus would, therefore, be beneficial. Heparin-bonded Carmeda Bioactive Surface (Medtronics Cardiopulmonary, Anaheim, Calif.) was evaluated in an in vitro model of extracorporeal circulation at standard-dose heparin (5 U/ml), to examine the effects of the surface treatment on activation of blood elements, and at reduced-dose heparin (1 U/ml), to determine whether surface-bound heparin would serve as an effective anticoagulant. During the initial recirculation period, platelet counts in the Carmeda (n ˆ 12) circuits were preserved at both doses of heparin and compared with control values (n ˆ 12). Furthermore, plasma levels of platelet factor 4 and beta-thromboglobulin were significantly reduced in the Carmeda circuits throughout the experiment. Moreover, as heparin concentration was reduced, the Carmeda surface treatment significantly decreased generation of C3a. Although heparin bonding was originally intended to obviate the need for systemic heparinisation, the Carmeda treatment did not reduce fibrinopeptide A generation at the lower dose of heparin. In summary, the Carmeda treatment failed to exhibit anticoagulant efficacy in this model; however, the data suggest that surface modification may have a role in ameliorating the typical inflammatory response initiated by blood contact with an artificial surface [64]. Heparin-coated extracorporeal circuits allow reduced amounts of systemic heparin and protamine. However, the effects on the coagulation and fibrinolytic systems when reducing systemic anticoagulation have partly remained unknown. In one study, 33 patients undergoing elective first time myocardial revascularisation were prospectively randomised either to have a cardiopulmonary bypass (CPB) circuit completely coated with covalently bound heparin, in combination with reduced systemic heparinisation (activated clotting time (ACT) > 250 s (n ˆ 17), or to a control group perfused with identical but uncoated circuits and full heparin dose (ACT > 480 s) (n ˆ 16)). Results showed

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that during CPB, the plasma level of prothrombin fragment 1.2 (PF 1.2) increased and was significantly higher than seen in the control group. However, the increase on CPB was modest compared to the major elevation observed after completed surgery and reversal of the anticoagulation. A similar pattern was observed for the thrombin-antithrombin III (TAT) complex. The platelet release of beta thromboglobulin increased in both groups during CPB and significantly more in the control group at the end of bypass (P < 0:01), indicating less platelet activation in the heparin-coated group. There were no significant intergroup differences with regard to fibrinolytic activity. Plasma fibrinogen as well as platelet counts were unchanged after the operation, compared to baseline. This study indicates that completely heparin-coated CPB can safely be performed in combination with reduced systemic heparinisation. The heparin and protamine amounts could be lowered to 35% of normal doses. Indications of more thrombin generation on CPB compared to the uncoated controls were seen, but the levels remained within low ranges in both groups. There was no evidence of thromboembolic episodes or clot formation in the extracorporeal circuits [65]. Clinical effect Heparinisation of the blood contact surface in cardiopulmonary bypass circuits has been promoted as an important step in the development of open heart surgery. As it decreases the inflammatory response resulting from the extracorporeal circulation, it may have a positive effect on clinical outcomes. This meta-analysis was carried out to examine if heparin-bonded circuits (HBCs) reduce the need for blood products and improve overall clinical outcome. A systematic literature search was performed to identify randomised controlled trials reporting outcomes of HBCs compared with non-HBCs. Primary outcomes assessed were postoperative blood/blood-product transfusion and blood loss. Secondary outcomes included all-cause mortality, acute postoperative myocardial infarction, stroke, re-sternotomy for post-operative bleeding, wound infection, atrial fibrillation, duration of ventilation, intensive care unit (ICU) and hospital-length of stay (LOS). Random effects meta-analytical techniques were applied to identify differences in outcomes between the two groups. Quality of the included studies and heterogeneity were assessed. From an initial review of 762 published studies, 41 randomised trials fulfilled the inclusion criteria, leaving 3434 patients' data for analysis. HBCs significantly decreased the incidence of blood transfusion required. It also significantly decreased resternotomy, duration of ventilation, ICU-LOS and hospital-LOS. HBCs had no effect on other adverse events evaluated. Although HBCs showed a positive effect on some of the clinical outcomes, it showed only marginal differences for other outcomes. Further evaluation of the cost-effectiveness of this technology is required [66].

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Summary

Inspired by the endothelium non-thromobogenic surface, which is able to release anti-thrombogenic substances to inhibit blood clotting process, heparinisation has been well developed for improving blood compatibility of biomaterials. Surface heparinisation can be achieved by many approaches such as physical, covalent, crosslinking, electrostatic and entrapment methods. The costs for modification are variable, while the covalent method is the highest and the physical method is the lowest. Physical methods include coating via ionic bonding. Covalent bonding normally involves significant chemical treatments both for materials and heparin sometimes. Heparinisation can significantly reduce platelet adhesion and inhibit blood coagulation. In clinics, surface heparinisation of blood contacting devices is mainly to reduce blood loss and inflammatory response, to minimise the implication of the usage of heparin, and to improve blood compatibility. Although heparinisation has some effects on inhibiting blood clotting, heparin coating sometimes is limited in its effectiveness and its applicability to varied uses [67]. This promotes the future exploring new methods for improving blood compatibility of biomaterials.

3.5

Future trends in heparinisation of biomaterial surfaces

For improving of blood compatibility of biomaterial, surface heparinisation is one of the most effective ways in the medical device industry. Since the blood response of the surface is not only dependent on the surface immobilisation of heparin, the configuration of the heparin molecules, the flexibility of immobilised heparin, the combination of heparin with other molecules, the surface charges and the morphology are also important. For example, by utilisation of a spacer like PEO, albumin to form heparin conjugates and then the conjugates are to be coated onto a polymer surface, the surface will be more blood compatible due to the introduction of the flexible layer [68]. The influence of heparinisation on antibacterial properties is also an important area as device infection is always a concern when implants are used. The coating of the surface with endothelial cells is a very attractive method to solve the issue of blood compatibility of biomaterials [69]. The clinical challenges in the heparinised biomaterials in medical devices are still in the aspects of the longevity, cost, efficacy and regulation control. The cost and efficacy need to be balanced as the most sophisticated surface heparinisation is sometimes difficult to use clinically. Tissue engineering and stem cell used to regenerate blood vessels are future perspectives, while the polymer scaffolding surface might need heparinisation to regulate angiogenesis.

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Developing the vascular network in tissue engineering is one of the challenges today. To regulate the process of angiogenesis requires the interaction between cells and certain regulators, such as growth factors. Some heparinbinding growth factors such as basic Fibroblast Growth Factor (bFGF) and Vascular Endothelial Growth Factor (VEGF or VEGF-A) are the growth factors affecting this process. By surface modification of the scaffolding materials with heparin, these growth factors can be bound to the heparinised materials such as collagen matrix [70], calcium phosphate bone cements [71] and some porous PLGA biodegradable scaffolds [72] to achieve tissue regeneration. The surface heparinisation approach can be one of the simple methods to introduce growth factors into the biomaterial matrix for vascular tissue engineering.

3.6

References

[1] MX: Business Strategies for Medical Technology Executives, in its May/June 2008 edition. [2] Ramakrishna S, Mayer J, Wintermantel E, Leong KW (2001) Biomedical applications of polymer-composite materials: a review, Composites Science and Technology 61: 1189±224. [3] Lister J (1863) On the coagulation of blood. Proc R Soc London 12: 580. È ber die Ursache der Blutgerinnung. Med Jahrb Wien 3: 259. [4] Freund E (1885) U [5] Bordet J, Gengou O (1903) Recherches sur la coagulation du sang. Ann de 1'Inst Pasteur 17: 822. [6] Zhao X (1999) Blood response of plasticised poly(vinyl chloride): influence of surface modification, PhD Thesis, University of Strathclyde. [7] Neubauer O, Lampert H (1930) Die physikalische Seite des Blutgerinnungsproblems. Muench Med Wochenschr 77: 582. [8] Sawyer PN, Burrowes C, Ogoniak J, Smith AO, Wesolowski SA (1953) Ionic architecture at the vascular wall interface. Trans Am Soc Artif Intern Organs 10: 316±19. [9] Gott VI, Whiffen JD, Dutten RC (1963) Heparin bonding colloidal graphite surfaces. Science 142: 1297. [10] Zisman WA (1964) Contact angle, wettability and adhesion, Fowkes FM (ed) Adv Chem Ser 43: 1±51. [11] Lyman DJ (1975) Polymers in medicine and surgery. In Kronenthal RL (ed) Polymer Science and Technology, Vol. 8. Plenum Press, New York [12] Hecher JF, Edwards RO (1981) Effects of roughness on the thrombogenicity of a plastic. J Biomed Mater Res 15: 1±7. [13] Hayward JA, Chapman D (1984) Biomembrane surfaces as models for polymer design: the potential for haemocompatibility. Biomaterials 5: 135±42. [14] Ishihara K, Oshida H, Endo Y, Ueda T, Watanabe A, Nakabayashi N (1992) Hemocompatibility of human whole blood on polymers with a phospholipid polar group and its mechanism. J Biomed Mater Res 26: 1543±52. [15] Shulman RI, Singer M, Rock J (2002) Keeping the circuit open: lessons from the lab. Blood Purificat 20: 275±81. [16] McLean J (1916) The thromboplastics action of cephalin. Am J Physiol 41: 250. [17] Abramson S, Niles JL (1999) Anticoagulation in continuous renal replacement therapy. Curr Opin Nephrol Hypertens 8: 701±7.

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[18] Clowes A, Karnovsky M (1977) Suppression by heparin of smooth muscle cell proliferation in injured arteries. Nature 265: 625. [19] Sy M, Schneeberger E, McCluskey R, Greene M, Rosenberg R, Benacerraf B (1983) Inhibition of delayed-type hypersensitivity by heparin depleted of anticoagulant activity. Cellular Immunology 82: 23±8. [20] Thomas K, Gimenez-Gallego G (1986) Fibroblast growth factors: broad spectrum mitogens with potent angiogenic activity. Trends Biochem Sci 11: 81. [21] Damus PS, Hicks M, Rosenberg RD (1973) Anticoagulant action of heparin. Nature 246: 355. [22] Anderson JM, Bonfield TL, Ziats NP (1990) Protein adsorption and cellular adhesion and activation on biomedical polymers. Int J Arif Organs 13: 375±82. [23] Vroman L (1988) The life of an artificial device in contact with blood: Initial events and their effect on its final state. Bull NY Acad Med 64: 352±7. [24] Andrade JD, Hlady V (1986) Protein adsorption and material biocompatibility. A tutorial review and suggested hypotheses. Adv Polym Sci 79: 1±63. [25] Forbes CD, Courtney JM (1994) Thrombosis and artificial surfaces. In: Bloom AL, Thomas DP, Forbes CD, Tuddenham EGD (eds) Haemostasis and Thrombosis. Edinburgh, Churchill Livingstone, pp 1301±24. [26] Yin HQ, Zhao XB, Courtney JM, Blass CR, West RH, Lowe GDO (1999) Blood intereactions with plasticised poly(vinyl chloride): relevance of plasticiser selection. J Mater Sci: Mater Med 10: 527±31. [27] Courtney JM, Lamba NMK, Sundaram S, Forbes CD (1994) Biomaterials for blood-contacting applications. Biomaterials 15: 737±44. [28] Buck RG, Scarborough DE, Saba SR, Brinkhous KM, Ikenberry LD, Kearney JJ, Clark HG (1969) Thrombogenicity of some biomedical materials: platelet±interface reactions. J Biomed Mater Res 3: 615±44. [29] Mohandas N, Hochmuth RM, Spaeth EE (1974) Adhesion of red cells to foreign surfaces in the presence of flow. J Biomed Mater Res 8: 119±36. [30] Kazatchkine MD, Carreno MP (1988) Activation of the complement system at the interface between blood and artificial surfaces. Biomaterials 9: 30±5. [31] Diamantoglou M, Vienken J (1996) Strategies for the development of haemocompatible dialysis membranes. Macromol Symp 103: 31±42. [32] Piskin E (1992) Biologically Modified Polymeric Surfaces. Amsterdam, Elsevier Applied Sciences. [33] Andelman B, Rizk A, Hanners E (1988) Plasminogen interactions with platelets in plasma. Blood 72: 1530±5. [34] Miller KM, Anderson JM (1988) Human monocyte/macrophage activation and interleukin, I. Generation by biomedical polymers. J Biomed Mater Res 22: 713±31. [35] Grode GA, Andersson SJ, Grotta HM, Falb RD (1969) Nonthrombogenic materials via a simple coating process. Trans Am Soc Artif Intern Organs 15: 1. [36] Merker RL, Elyash LJ, Mayhew SH, Wang JYC (1969) In Hegyeli RJ (ed) Artificial Heart Program Conference Proceedings, Washington, DC, National Heart Institute Artificial Heart Program, 29. [37] Britton RA, Merrill EW, Gilliand ER, Salzman EW, Austen WG, Kemp DS (1968) Antithrombogenic cellulose film. J Biomed Mater Res 2: 429. [38] Lagergren MR, Eriksson JC (1971) Plastics with a stable surface monolayer of crosslinked heparin. Trans Am Soc Artif Intern Organs 17: 10. [39] Heyman PW, Cho CS, McRea JC, Olsen DB, Kim SW (1985) Heparinised polyurethane: in vitro and in vivo studies. J Biomed Mat Res 19: 419±36. [40] Raghunath K, Biswas K, Rao KP, Joseph KT, Chvapil M (1983) Some

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characteristics of collagen-heparin complex. J Biomed Mat Res 17: 613±21. [41] Byun Y, Jacobs, HA, Kim SW (1995) Heparin surface immobilization through hydrophilic spacers: thrombin and antithrombin III binding kinetics. J Biomater Sci, Polym Ed 6: 1±13. [42] Han DK, Park KD, Ahn KD, Jeong SY, Kim YH (1989) Preparation and surface characterization of PEO-grafted and heparin-immobilized polyurethanes. J Biomed Mater Res 23(A1 Suppl): 87±104. [43] Nagaoka S, Kurumatani H, Mori Y, Tanazawa H (1989) The activity of immobilized heparin via long poly(ethylene oxide) spacers. J Bioactive Compat Polym 4: 323±32. [44] Magoshi T, Matsuda T (2002) Formation of polymerized mixed heparin/albumin surface layer and cellular adhesional responses. Biomacromolecules 3(5): 976±83. [45] Michanetzis GPA, Katsala N, Missirlis YF (2003) Comparison of haemocompatibility improvement of four polymeric biomaterials by two heparinization techniques. Biomaterials 24: 677±88. [46] Bamford CH, AI-Lamee KG (1996) Studies in polymer surface modification and grafting for biomedical uses. Polymer 22: 4885±9. [47] Bamford CH, Al-Lamee KG (1994) Studies in polymer surface functionalization and grafting for biomedical and other applications. Polymer 13: 2844±52. [48] Seifert B, Groth T (1995) Immobilization of heparin on polylactide for application to degradable biomaterials in contact with blood. J Biomater Sci Polym Ed 7(3): 277±87. [49] Mino G, Kaizerman S (1958) The new method for the preparation of graft copolymers. Polymerization initiated by ceric ion redox systems. J Polym Sci 31(122): 242±3. [50] Peppas NA, Merill EW (1977) Development of semicrystalline PVA networks for biomedical applications. J Biomed Mater Res 11: 423±34. [51] Bae J-S, Seo E-J, Kang I-K (1999) Synthesis and characterization of heparinized polyurethanes using plasma glow discharge. Biomaterials 20: 529±37. [52] http://www.carmeda.se/upl/files/1183.pdf [53] Zhao X, Courtney JM, Yin HQ, West RH, Lowe GDO (2008) Blood interactions with plasticised poly(vinyl chloride): influence of surface modification. J Mater Sci: Mater Med 19: 713±19. [54] Liu X-Y, Zhang C-C, Xu W-L, Ouyang C-X (2009) Controlled release of heparin from blended polyurethane and silk fibroin film. Mater Lett 63(2): 263±5. [55] Farndale RW, Buttle DJ, Barret AJ (1986) Improved quantitation and discrimination of sulphate glycosaminoglycans by use of dimethylmethylene blue. Biochim Biophys Acta 883: 173±7. [56] Van de Lest CH, Versteeg EM, Veerkamp JH, van Kuppevelt TH (1994) A spectrophotometric method for the determination of heparin sulfate. Biochim Biophys Acta 1201: 305±11. [57] Ito Y, Sisido M, Imanishi Y (1986) Synthesis and antithrombogenicity of anionic polyurethanes and heparin-bound polyurethanes. J Biomed Mater Res 20: 1157±77. [58] Kang I-K, Kwon OH, Lee YM, Sung YK (1996) Preparation and surface characterization of functional group-grafted and heparin immobilized polyurethanes prepared by plasma glow discharge. Biomaterials 17: 841±7. [59] Sanchez J, Elgue G, Riesenfeld J, Olsson P (1998) Studies of adsorption, activation and inhibition of factor XII on immobilized heparin. Thromb Res 89: 41±50. [60] Keuren JF, Wielders SJ, Driessen A, Verhoeven M, Hendriks M, Lindhout T (2004) Covalently-bound heparin makes collagen thromboresistant. Arterioscler Thromb

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Vasc Biol 24(3): 613±17. [61] Keuren JF, Wielders SJ, Willems GM, Morra M, Cahalan L, Cahalan P, Lindhout T (2003) Thrombogenicity of polysaccharide-coated surfaces. Biomaterials 24(11): 1917±24. [62] Meinhardt JP, Annich GM, Miskulin J, Kawai T, Ashton BA, Bartlett RH (2003) Thrombogenicity is not reduced when heparin and phospholipid bonded circuits are used in a rabbit model of extracorporeal circulation. ASAIO J 49(4): 395±400. [63] Niimi Y, Ichinose F, Ishiguro Y, Terui K, Uezono S, Morita S, Yamane S (1999) The effects of heparin coating of oxygenator fibers on platelet adhesion and protein adsorption. Anesth Analg 89(3): 573±9. [64] Korn RL, Fisher CA, Livingston ER, Stenach N, Fishman SJ, Jeevanadam V, Addonizio VP (1996) The effects of Carmeda bioactive surface on human blood components during simulated extracorporeal circulation. J Thorac Cardiovasc Surg 111(5): 1073±84. [65] Ovrum E, Brosstad F, Am Holen E, Tangen G, Abdelnoor M, Oystese R (1996) Complete heparin-coated (CBAS) cardiopulmonary bypass and reduced systemic heparin dose; effects on coagulation and fibrinolysis. Eur J Cardiothorac Surg 10(6): 449±55. [66] Mangoush O, Purkayastha S, Haj-Yahia S, Kinross J, Hayward M, Bartolozzi F, Darzi A, Athanasiou T (2007) Heparin-bonded circuits versus nonheparin-bonded circuits: an evaluation of their effect on clinical outcomes. Eur J Cardiothorac Surg 31(6): 1058±69. [67] Stratta P, Canavese C, Mangiarotti G (1981) Heparin is unable to prevent activation by three different membranes. Proc. EDTA 18: 269. [68] Liu L, Ito Y, Imanishi Y (1991) Synthesis and antithrombogenicity of heparinised polyurethanes with intervening spacer chains of various kinds. Biomaterials 12(4): 390. [69] Miwa H, Matsuda T, Kondo K (1992) Improved patency on an elastomeric vascular grafts by hybridisation. ASAIO J 38(3): 512. [70] Yao C, PreÂvel P, Koch S, Schenck P, Noah EM, Pallua N, Steffens G (2004) Modification of collagen matrices for enhancing angiogenesis. Cells Tissues Organs 178: 189±196. [71] Lode A, Reinstorf A, Bernhardt A, Wolf-Brandstetter C, Konig U, Gelinsky M (2008) Heparin modification of calcium phosphate bone cements for VEGF functionalization. J Biomed Mater Res 86A: 749±59. [72] Yoon JJ, Chung HJ, Lee HJ, Park TG (2006) Heparin-immobilized biodegradable scaffolds for local and sustained release of angiogenic growth factor. J Biomed Mater Res 79A: 934±42.

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Surface modification of biomaterials by peptide functionalisation L . S . B I R C H A L L , H . Q U and R . V . U L I J N , University of Strathclyde, UK

Abstract: Surface science provides crucial tools for the design of nextgeneration biomaterials, as it is at the surface where interactions between man-made materials and the biological environment occur. Peptides constitute biology's molecular language and play key roles in the development of surfaces that can detect, respond to and ultimately direct biomolecular events at the solid±liquid interface. This chapter reports on the production of surfaces that are functionalised with peptides using a variety of strategies: covalent peptide functionalisation strategies using step-wise and one-step methods, as well as non-covalent molecular self-assembly at flat surfaces and in three dimensions to form functional nanostructures. Key words: peptide, surface, biomaterial, solid-phase, bioactive.

4.1

Introduction

Surface science provides crucial tools for design of next-generation biomaterials as it is at the surface where interactions between man-made materials and the biological environment occur. Biomolecules such as polypeptides, nucleotides and saccharides constitute biology's molecular language and can therefore play key roles in development of surfaces that can detect, respond to and ultimately direct biomolecular events at the solid-liquid interface. Here, we will focus on the production and analysis of surfaces that are functionalised with oligopeptides (Branco et al., 2009). We will discuss covalent peptide functionalisation strategies using step-wise and one-step functionalisation methods, as well as non-covalent molecular self-assembly at flat surfaces (2D) and in three dimensions (3D). Top-down and bottom-up approaches will be discussed that may be used to achieve spatially resolved presentation of peptides for application in designed molecular biomaterials.

4.2

Peptides and peptide functionalisation of surfaces

Peptides are short polymers comprising up to 50 amino acids that are linked together via amide bonds. They are composed of 20 gene encoded amino acids

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(plus a number of post-translational modified and unnatural ones) that differ in their side chain chemistry, which may be aromatic, aliphatic, polar, and/or charged (or combinations thereof). Peptides therefore provide an extremely rich and versatile chemical toolbox. Amino acids are often abbreviated to one or three letter codes, with peptide sequences written in the direction from N- to Cterminus. The one letter code is often used to concisely represent long peptide sequences (Fig. 4.1). In this chapter we will use the one letter code to describe peptide sequences. Biology provides numerous examples where peptides (usually in the context of large proteins) are responsible for communication between cells and their environments. This observation suggests that peptide functionalised surfaces may provide a key route to engineering an effective interface between biological systems and man-made materials. Indeed, peptide functionalised surfaces have been extensively utilised in biomaterials science (Chow et al., 2008; Mart et al., 2006). Many groups have exploited their properties to form molecular architectures for a variety of different applications (Yang et al., 2008; Zhang, 2003; Palmer and Stupp, 2008; Estroff and Hamilton, 2004; Banwell et al., 2009; Lutolf and Hubbell, 2005; Reches and Gazit, 2006), including biosensing (Birchall et al., 2008; Laromaine et al., 2007; Welser et al., 2009; Boeneman et al., 2009; Medintz et al., 2006; Yao et al., 2007; Meldal, 2002), controlled release/drug delivery (McDonald et al., 2009; Thornton et al., 2008), and designed biomaterials for cell culture and tissue engineering (Todd et al., 2007; Zhou et al., 2009; Dreher et al., 2008; Escuder and Miravet, 2006; Dalsin et al., 2003; Silva et al., 2004; Kisiday et al., 2002; Haines-Butterick et al., 2007). The above developments have been made possible by the discovery of solid phase peptide synthesis in 1963 by Merrifield, which has dramatically facilitated their synthesis. Broadly, three methods now exist for the incorporation of peptides into biomaterials: (i) building them up step-wise directly onto the material surface, (ii) grafting pre-synthesised peptides in one step, (iii) exploiting non-covalent self-assembly at 2D surfaces or in 3D. A discussion of protecting groups, coupling reagents, activation strategies and methods for peptide synthesis and chemical attachment will be given, highlighting their advantages and disadvantages. A number of strategies for peptide self-assembly allowing for dynamic presentation of peptides will also be discussed. In many applications, spatially resolved peptide functionality is important, for example in peptide biochips and in the design of biomaterials that support interfaces between different cell or tissue types. A number of surface fabrication techniques have been developed, which allow for the control of spatial resolution from the macroscale to the nanoscale. These techniques include photolithography, soft lithography, inkjet printing and dip pen nanolithography. In order to characterise the resulting modified surfaces, techniques such as AFM, SEM, TEM, UV, fluorescence, MALDI-ToF-MS, ToF-SIMS, XPS and

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4.1 The chemical structures of the 20 naturally occurring amino acids grouped according to the properties of their functional groups. Their three and one letter codes are given below each entry. The highlighted amino acids represent those that have both hydrophobic and aromatic characteristics.

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confocal/two- photon microscopy are typically employed of which some examples will be discussed (see Chapters 8 and 9).

4.3

Defining the biomaterial surface

Throughout this chapter biomaterial surfaces will be discussed in two contexts: two dimensional (2D) and three dimensional (3D). The distinction between 2D and 3D is rather arbitrary as it depends on the relative size/topography/curvature of surface features compared to that of the biological component interacting with it (Stevens and George, 2005). If the biological component is smaller (i.e., by approximately one order of magnitude) in comparison to the surface it interacts with, this surface is defined as 2D. If the biological component is similar in size or larger in contrast to the surface feature, the material is defined as 3D. For the purpose of this chapter we refer to nanotopographical surfaces and nanoparticles as 3D and microtopographical surfaces (and upwards) as 2D. Two approaches that will be discussed throughout this chapter are `top-down' and `bottom-up' methods for the production of functionalised materials. The top-down approach relies on microfabrication techniques, where externally controlled tools are used to cut and etch materials into the desired shape and order for their application. Such methods include micropatterning, photolithography and inkjet printing, which are especially suited to the production of 2D surfaces with micronscale features (although specialised techniques such as dip-pen nanolithography (Piner et al., 1999) and NSOM lithography (Reynolds et al., 2009) are available to achieve nanoscale resolution). The bottom-up approach utilises the physical and chemical properties of peptide molecules to form well-defined surfaces or molecular self-assembled systems with higher resolution obtainable, which is appropriate in the production of 3D structures.

4.4

Peptide functionalised surfaces

Functionality can be introduced to a surface via chemical synthesis or by molecular self-assembly (Ozin and Arsenault, 2005). The functionalisation strategy to be followed is usually dictated by chemical composition of the material rather than the scale of the surface (Fig. 4.2a). Peptide functionalisation can be carried out on metals (e.g. titanium or gold), glass, ceramics and polymers. By introducing chemical group functionality to these surfaces, peptides can be covalently attached utilising a variety of coupling reactions (see Section 4.4.1). Self-assembled monolayers (SAMs) of alkane thiols on gold have been extensively used to produce 2D surfaces on a variety of materials that can be further functionalised with biomolecules. Alternatively cysteine (with ±CH2± SH side chains) containing peptides have been used to directly anchor peptides to surfaces (as discussed in Section 4.4.2). Similarly, gold nanoparticles may be functionalised via attachment of thiol containing peptides.

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4.2 Covalent approach: the general strategy of SPPS on 2D/3D biomaterials surface. Where X ˆ N-terminal protecting group, Y ˆ amino acid side chain protecting group and R ˆ amino acid residue.

As described above, peptide functionalisation can be achieved either via stepwise synthesis (Fig. 4.2b) or using a one-step coupling process (Fig. 4.2c) of a pre-synthesised peptide. The advantage of step-wise synthesis to functionalise a surface is that the desired peptide is rapidly synthesised directly on the surface, any byproducts and unreacted starting materials can be washed away, meaning further purification is not required. There are two approaches utilised for step-

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wise synthesis on surfaces: photolithography and SPOT-synthesis, which will be discussed briefly in the next section. These methods can be easily automated and allow for the production of large libraries of peptides in parallel, using small quantities of reagents in a rapid manner.

4.4.1

Step-wise synthesis

Merrifield first described SPPS in 1963 (Merrifield, 1963), where he demonstrated the synthesis of a tetrapeptide on a solid support. This Nobel Prize winning advance in Chemistry is now the standard way of producing peptides and proteins in a synthetic manner. It has opened the door to the synthesis of natural proteins which are difficult to express in bacteria, the incorporation of unnatural amino acids, permits backbone modification and allows the synthesis of D-proteins and peptides (which contain D-amino acids). While traditionally performed on functionalised polymeric beads that have good swelling properties in organic solvents (such as polystyrene) researchers are increasingly performing SPPS directly onto biomaterial surfaces that allow for chemical functionalisation and biochemical assay to be performed on the same platform. Examples of materials that have been used in this context are cellulose (Frank, 1992), poly(ethyleneglycol acrylamide) co-polymer (PEGA) in the form of beads (Meldal, 1992; Ulijn et al., 2002) or (micropatterned) surfaces (Zourob et al., 2006; Todd et al., 2007), PEG-amine monolayers on glass (Todd et al., 2009), aminosilanised silicon oxide (Mosse et al., 2009; Rawsterne et al., 2006, 2007) and SAMs of alkane thiols on gold (Laurent et al., 2008). In some cases, peptide functionalised surfaces (such as PEGA, which is biocompatible) were directly used as platforms for liquid crystal based enzyme sensing (Birchall et al., 2008) and as switchable surfaces for enzymatically controlled cell attachment (Todd et al., 2007). Unlike natural production of peptides and proteins, SPPS proceeds in a Cterminal to N-terminal fashion. The general procedure relies on step-wise synthesis: repeated cycles of covalently coupling a single protected amino acid to the amine functionalised solid support, followed by deprotection of the Nterminus, allowing the next amino acid to be coupled. Once synthesis is complete the amino acid side chains can be deprotected and the peptide then cleaved from the solid support (or used further on-support as described above). The use of an insoluble but porous solid support means that the growing amino acid chain is covalently linked to the surface, whilst other reagents and byproducts can be washed away via filtration. There are two main strategies used for SPPS, t-butoxycarbonyl (Boc) and Fluorenylmethoxycarbonyl (Fmoc). The Boc strategy was adopted by Merrifield and is still used when synthesising non-natural peptides which are base sensitive. The Fmoc-strategy (Carpino and Han, 1972) is more widely used now for routine synthesis as the final peptide product can be cleaved under milder basic

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Table 4.1 Typical protecting groups for SPPS

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conditions than Boc. The fluorogenic and chromogenic nature of the fluorenylgroup provides a convenient spectroscopic handle for monitoring reactions. A plethora of protecting groups are used for side chain and N- and C-terminal protection (Table 4.1); such protecting group strategies can be selected to allow further functionalisation of the growing amino acid chain. The removal of these groups from the peptide chain and subsequent cleavage from the solid support can become complex and cleavage mixtures should be chosen with some thought, e.g. M, C and Y residues are susceptible to alkylation by cations produced in the cleavage process, adding scavengers to the cleavage mixture can largely suppress such side reactions. The reader is encouraged to investigate the various synthesis strategies available for the production of the desired peptide chain (Greene and Wuts, 1999; Valeur and Bradley, 2009; Montalbetti and Falque, 2005). In particular, the selection of protecting groups can prove very important in preventing unwanted reactions when deprotecting side chain residues. Chemical suppliers often provide invaluable information for the design of SPPS strategies. Various activating agents can be used to couple the amino acids to the amine residues on the solid support (Montalbetti and Falque, 2005; Valeur and Bradley, 2009). The common mechanisms rely on the activation of the carboxyl group. Carbodiimides (Table 4.2, entries 1 to 3) react with the carboxylic acid of the reagent to form the highly reactive O-acylurea, which then reacts directly with the solid-supported amine to yield the desired amide and a urea by-product. The O-acylurea can also lead to oxazolone formation which leads to epimerisation (racemisation) (Izumiya and Muraoka, 1969) whilst acetyl transfer to form the unreactive N-acylurea is often observed. The addition of benzotriazole reagents (e.g., Table 4.2, entries 4 and 5) to the reaction is typically used to overcome this problem. Recently, ethyl cyano(hydroxyimino) acetate (Table 4.2, entry 6) has been developed as a non-explosive alternative to HOBt. Currently, onium salts based upon the triazoles have been introduced for the efficient catalysis of amide bond formation (Albericio et al., 1998) (e.g., Table 4.2, entries 7 to 9).

4.4.2

One-step functionalisation

The alternative method to step-wise synthesis relies on the direct immobilisation of a pre-synthesised peptide onto the (biomaterial) surface. Here, the peptide is first synthesised on a polystyrene solid support via step-wise synthesis and then cleaved and purified before attaching it to the required surface. With this approach, the amino acid side chain residues (e.g., ±NH2 of lysine or the ±SH of cysteine) are often used as a mode of attachment, but frequently specific functionality must be built into the molecule prior to immobilisation. A variety of coupling reactions have been used for this purpose, using aqueous media or organic solvents. Table 4.3 shows a number of these, detailing the functional

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Table 4.2 Common activating agents used in SPPS

groups utilised on the surface (the substrate) and the chemistry used to attach the peptides (the reagents). Peptides with amine, alcohol, thiol or acid functionalities can all be attached to epoxide functionalised surfaces without the need for additional reagents and under dry conditions (Table 4.3, entries 1± 4). Amide bond formation (i.e., condensation reaction between an amine and a carboxylic acid) can be carried out using a variety of activating agents (see above) (Table 4.3, entry 5). If the carboxylic acid is tethered to the surface it is first activated, e.g. by conversion to the N-hydroxysuccinimide ester, before coupling of the peptide to a free primary amine (Table 4.3, entry 11). 4-(pmaleimideophenyl)butyric acid N-hydroxysuccinimide ester (SMPB) is covalently bound to amine surfaces through the hydroxysuccinimide (Table 4.3, entry 6). The resulting maleimide functionality can then be used to attach

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Table 4.3 Functional groups utilised in the covalent attachment of peptides to surfaces

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Table 4.3 Continued

further peptides (Table 4.3, entry 9). Alkoxyamine containing peptides can be tethered to aldehyde functionalized surfaces via oxime formation (Table 4.3, entry 7). Amino-thiol containing peptides can be tethered to glyoxylyl functionalized surfaces via thiazolidine formation (Table 4.3, entry 8), whilst peptides with a terminal cysteine residue can be attached directly to surfaces with disulfide functionality via thiol-disulfide exchange, in this instance the control of oxidation must be considered (Table 4.3, entry 10).

4.5

Non-covalent peptide functionalisation by selfassembly

Molecular self-assembly has become a powerful tool for fabricating 2D surfaces (SAMs) as well as supramolecular 3D architectures. Oligopeptide SAMs can be created by incorporation of cysteine residues, providing the ±SH groups, that allow them to be anchored onto gold surfaces (Zhang et al., 1999) or nanoparticles (Laromaine et al., 2007; Wang et al., 2005) via gold-thiol bond formation. The opposite end of the peptide may carry a variety of functionalities to interface directly with biological systems. Self-assembly of three-dimensional structures is mediated by weak and noncovalent bonds such as hydrogen bonds, electrostatic interactions, hydrophobic interactions, -stacking and Van der Waals interactions (Zhang, 2003). Peptides and their derivatives are versatile building blocks in this regard and the design rules for peptide building blocks are emerging. These design rules are either

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derived by copying nature (-helix, -sheet) or may exploit entirely new designs based on peptide derivatives (aromatic and aliphatic amphiphiles) (Ulijn and Smith, 2008). Many of these structures form fibres which may become entangled to form hydrogel networks. A number of approaches have been developed to control self-assembly (gelation) in an on-demand fashion (by exploiting external stimuli such as pH, light or enzyme action) (Mart et al., 2006). The resulting materials allow for precise presentation of peptides on the material surface, which may lead to a number of applications in biotechnology (3D cell culture, biosensing) and technology (nanoelectronics and templating). Shuguang Zhang first demonstrated the use of -sheet forming peptides to form nanostructures in the early 1990s (Zhang et al., 1993). -sheets consist of multiple peptide chains that have an extended backbone arrangement that permits hydrogen bonding between the backbone amides and carbonyl groups. -sheets can be orientated so that all their C-termini are at one end of a structure, described as a parallel structure, or so that the N- and C-termini alternate, described as an anti-parallel structure. This has an important impact on the orientation of hydrogen bonds between sheets and on the stability of the resulting structure. There is a basic motif present in most -sheets which consists of alternating hydrophobic, hydrophilic residues (Fig. 4.3a). A number of approaches have utilised ionic self-complementary peptides (e.g., containing RADA repeat units) which have been used to form well-defined nanofibre scaffolds (Yokoi et al., 2005). These sixteen residue peptides form stable sheets via electrostatic interactions, which undergo molecular self-assembly into nanofibres and eventually a scaffold hydrogel. In these systems functional peptides can be introduced by end-functionalisation of a proportion of the peptide building blocks, which has been exploited in production of a range of bioactive gels that direct (stem) cell behaviour (Zhao and Zhang, 2006). An alternative to -sheets are -hairpins, consisting of two short -sheet sequences linked by a turn sequence, making it possible for peptides to stack in register with one another. One designed system by Schneider and Pochan, uses a turn sequence ±VDPPT± flanked by two eight amino acid sequences with a high -sheet propensity. These materials have been studied in the context of 3D cell culture (Haines-Butterick et al., 2007) and controlled release of model drugs (Branco et al., 2009). A final naturally derived structural motif that has been exploited to produce fibrous structures and hydrogels by molecular self assembly are -helices, commonly used as components of coiled-coils. The design rules for these systems are now well understood, and are based on repeats of seven residues (a heptad) (Woolfson and Ryadnov, 2006). A recent example of the use of coiledcoiled systems included formation of hydrogels for cell culture (Banwell et al., 2009). Peptide amphiphiles (PA), composed of peptides appended with aliphatic tails, have been designed to reversibly self-assemble into nanofibrous networks,

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4.3 (a) RADA repeat sequences self-assembly via electrostatic interactions, (b) Peptide amphiphiles self-assembly into nanofibrous networks, (c) Fmocpeptides self-assembly via ± interactions for nanocylinders.

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through changes in pH or ionic composition (Fig. 4.3b) (Hartgerink et al., 2002). These fibres can reversibly polymerise to enhance their stability via the formation of intermolecular disulfide bridges. Bioactive peptides, PAs including RGD (Mata et al., 2009) and IKVAV (Silva et al., 2004), have also been presented on the surface of fibres in very high density and have found applications in 3D cell culture. A relatively new class of self-assembling systems that use much shorter peptide sequences are aromatic peptide amphiphiles: peptides appended at the N-terminus with aromatic moieties such as fluorenyl, naphthalene, carboxybenzyl, and pyrene (Mahler et al., 2006; Jayawarna et al., 2006; Adams et al., 2009; Zhang et al., 2003). Interestingly, Fmoc, a common protecting group in peptide synthesis (see Section 4.4.1) has proved to be an excellent director of self-assembly in these systems, providing a straightforward means to impart self-assembling properties upon peptides simply by leaving the N-terminal protecting group attached. We recently proposed a novel supramolecular structure for this class of peptides composed of -interlocked -sheets (Smith et al., 2008; Williams et al., 2009). In this - peptide configuration (Fig. 4.3c) (Smith et al., 2008), H-bonding and aromatic interactions both contribute significantly, and, depending on the nature of the peptide, may give rise to formation of fibres (Jayawarna et al., 2006), sheets (Williams et al., 2009) or tubular structures (Das et al., 2008). Peptide fibres produced by this method inherently position peptide sequences at the surface. This feature has been exploited in the production of biomimetic nanofibrous hydrogel as a 3D-scaffold for anchorage-dependent cells. The peptide-based bioactive hydrogel is formed through molecular self-assembly and the building blocks are a mixture of two aromatic short peptide derivatives: Fmoc-FF and Fmoc-RGD, providing a highly hydrated, stiff and nanofibrous hydrogel network that uniquely presents bioactive ligands at the fibre surface in tunable density. This material was observed to promote adhesion of encapsulated dermal fibroblasts through specific RGD±integrin binding, with subsequent cell spreading and proliferation. Triggered self-assembly of aromatic peptide amphiphiles has been achieved using changes in pH (Jayawarna et al., 2006; Adams et al., 2009), dilution from organic solvent (Mahler et al., 2006) and by exploiting enzyme catalysed reactions (Toledano et al., 2006; Yang et al., 2004).

4.6

Spatial control of peptide functionality

Owing to the large numbers involved in peptide screening experiments, on-chip applications are increasingly employed, with automated synthesis and screening protocols considered to be highly attractive. Merrifield's (1963) SPPS is compatible with both manual and automated peptide synthesis, providing the opportunity to produce large arrays of peptides on solid supports for automated, rapid and parallel screening. This section will discuss ways of achieving spatial

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control of peptide functionality using a number of lithographic and printing approaches.

4.6.1

Photolithography

Peptide synthesis by photolithography on surfaces was first reported by Fodor et al. (1991). It is described as light-directed spatially controlled parallel chemical synthesis. The substrate surface is initially functionalised with amino groups which are blocked with a photolabile protecting group. Photo-deprotection is achieved by illumination of the substrate through a lithographic mask. The exposed amino groups are then accessible for the next coupling procedure via SPPS. Solid-phase chemistry, photolabile protecting groups and photolithography had been successfully combined to achieve an array of 1024 peptides on 1.6 cm2 amino derivatised substrate in 10 steps using 6-nitroveratroyloxycarbonyl (Nvoc)-protected (photolabile protecting group) peptides. This photolithographic peptide synthesis method can create high density synthesis sites, which allows 40 000 compounds to be synthesised in 1 cm2. Fodor and colleagues later developed photolithographic techniques to carry out parallel synthesis of large numbers of oligonucleotides on a solid surface, an approach that has been commercialised by Affymetrix for high-density DNA arrays (Heller, 2002).

4.6.2

SPOT-synthesis

Since its introduction many researchers have developed techniques to generate peptide sequences in biological assay systems in order to identify biomolecules and investigate biomolecular interactions. Frank (1992) first reported SPOTsynthesis for the generation of peptide sequences on cellulose. This technique has been developed for the preparation of a series of peptide sequences at the 50 nmol scale, utilising SPPS strategies. The spot size (typically 0.1±22 mm in diameter) is determined by the volume dispensed, the absorptive properties of the surface material, and the volatility of the solvent. It opened the door to the production of small macro-array surfaces. SPOT-synthesis was initially carried out manually on planar cellulose or laboratory filter paper with approximately 60-times larger diameters than microarrays spots (Blackwell, 2006). These increased the compound loading and eased compound quantification. Recently, Laurent et al. (2008) have reported peptide arrays, which are prepared by SPOTsynthesis, on SAM-coated gold surfaces for on-chip monitoring of the synthesis and subsequent evaluation of enzymatic modification of peptides by using MALDI-ToF mass spectrometry. SPOT-synthesis is a low cost, flexible, simple and rapid technique for parallel synthesis of large numbers of peptides and peptide mixtures. Furthermore, the peptide arrays produced by SPOT-synthesis have been used for a range of assay and screening methods such as epitope

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analysis, protein interactions and enzyme-substrate recognition (Reineke et al., 2001).

4.6.3

Soft lithography

Whitesides developed the process of soft lithography based on the SAM methodology (Ozin and Arsenault, 2005). It makes use of patterned elastomers, such as poly(dimethylsiloxane) (PDMS), as a mask, stamp or mould to print SAM patterns on surfaces. This area has been extensively reviewed elsewhere (Ozin and Arsenault, 2005). This soft lithographic technique has been further developed for micropatterning SAMs using microcontact printing. Microcontact printing utilises molecular self-assembly, that is molecules spontaneously assemble under thermodynamic equilibrium conditions, to form ordered structures. Zhang et al. (1999) combined self-assembling oligopeptide monolayers and microcontact printing for engineering cell patterned biological surfaces. Incorporation of a cysteine residue into the molecules deposited enabled anchoring to the surface via gold±thiol bond formation. It showed that oligopeptides could be well engineered to incorporate multiple features. Also, by designing the structural conformation of oligopeptides, tailoring additional bioactive ligands, linkers and anchors, and combining a protein resistant substrate, cells were accurately aligned and specific cell arrays were patterned efficiently. Soft lithography has also been exploited to pattern 3D self assembled structures, thereby combining bottom-up and top-down fabrication. Stupp et al. demonstrated self-assembly of RGD functionalized aliphatic peptide amphiphiles which were flow-aligned and micropatterned using moulds. UV polymerization of these materials gave stable gels with microscale topographical features that could be used to induce differentiation and provide directional control of mesenchymal stem cells (Mata et al., 2009).

4.6.4

Inkjet printing

In recent years, inkjet printing has been considered increasingly as a tool for the controlled deposition of polymers and functional material coatings for a number of applications, including micro-patterns for microelectronics, displays, biomedical devices and photonics. Compared with photolithographic and soft lithographic techniques, inkjet printing is inexpensive, repeatable, flexible and easy to use (Zhang et al., 2008; Liberski et al., 2009). The unique feature of inkjet printing is that it is a non-contact process. The droplets of ink are jetted from a small aperture or nozzle in the print-head and deposited onto a specified position on the substrate to create an image. The method by which the droplets are generated depends upon the particular technology used. The most common droplet generation methods of inkjet printing are the Continuous Inkjet (CIJ) and Drop-on-Demand (DOD)

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processes. To date, few examples exist of exploiting this method in peptide functionalisation. Gazit and co-workers have demonstrated the patterning of aromatic dipeptides nanotubes (Fmoc-FF, H-FF-OH, Boc-FF) (which had previously been used for fabrication of sensitive electrochemical biosensors) using a commercial inkjet printer (Adler-Abramovich and Gazit, 2008).

4.6.5

Dip-pen nanolithography (DPN)

Dip-pen nanolithography (DPN) was pioneered by the group of Mirkin (Piner et al., 1999; Salaita et al., 2007). They demonstrated that alkanethiols could be transferred from an AFM tip to a gold thin film. In general, a solution of molecules, the `ink', is transferred from an AFM tip, the `pen', to a solid state surface, the `paper', by capillary action (Fig. 4.4). This direct-write constructive lithography, allows the printing of both soft and hard materials onto surfaces with sub 50 nm resolution. Multiple compounds can be deposited on a surface sequentially or in parallel in a precise and predetermined manner. Most developments in this area have employed the deposition of larger molecules, such as proteins, DNA and even viruses. However, advances have been made in the patterning of peptides with a few examples discussed below. Collagen-like peptides have been printed onto gold surfaces at line widths of 30±40 nm and lengths of 100 nm. The peptide contained the primary amino acid sequence of native collagen, conserving the triple helical domain. The Nterminus of the peptide contained a cysteine residue for chemisorption to the gold surface (Wilson et al., 2001). Stupp's PA molecules have been delivered to silicon surfaces to produce closely packed and reproducible arrays of PA nanofibre patterns which measured approximately 6 nm in diameter. They were also able to produce

4.4 Schematic representation of dip-pen nanolithography.

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more complex patterns of nanostructures by carrying out multiple passes of the dip-pen (Jiang and Stupp, 2005). Coating the surface of the AFM tip with polyethylene glycol suppressed the assembly of nanofibres on the AFM tip. TAT peptides are derived from the TAT protein of HIV-type-1, which has been linked to important functions in cell permeation and viral replication. As it has similar recognition properties to that of the whole native protein it is frequently used for biological research. SMPB functionalised SiOx surfaces have been fabricated with TAT peptides using DPN. The terminal cysteine residue of the peptide was covalently attached to the surface by reacting with the maleimide functional group presented at the surface. The smallest feature they were able to write had a width of 100 nm (Cho and Ivanisevic, 2004). Similarly TAT peptides have been immobilised on gold surfaces. In contrast to the above approach, the peptide could be directly attached to the gold surface via the cysteine residue, without the need for pre-modification of the gold surface, the average width of the lines was around 100 nm (Cho and Ivanisevic, 2005).

4.7

Conclusions

It is widely recognised that peptides form a suitable interface between biology and man-made materials as peptides are synthetically accessible and are biology's expression language. Recent years have seen an explosion in the development of peptide functionalisation techniques of 2D and 3D surfaces using both covalent approaches and self-assembly. It is now possible to synthesise peptides directly onto biomaterial surfaces using either step-wise or one-step functionalisation strategies. Increasingly, the focus in this area is on achieving high spatial resolution and multifunctionality, in an objective to more closely mimic the anisotropy and multifunctionality observed in biological systems. Advanced patterning technologies such as DPN allow for high precision deposition; however, it is still challenging to produce multiple peptide functionalities in close proximity at nanoscale resolution. Another important new direction is to achieve dynamic control over peptide presentation, as observed in real biological systems. Progress in this direction has been achieved using stimuli responsive systems that are addressable by light (Kloxin, 2009), enzymes (Todd et al., 2009) and mechanical stimuli. While significant progress has been made, scientists are still far off achieving similar complexity to what is achieved by biological systems. Continued close collaboration between material scientists, chemists, physicists and biologists is essential for further progress in this area.

4.8

References

Adams, D. J., Butler, M. F., Frith, W. J., Kirkland, M., Mullen, L. and Sanderson, P. (2009) A new method for maintaining homogeneity during liquid-hydrogel transitions using low molecular weight hydrogelators. Soft Matter, 5, 1856±1862.

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Adler-Abramovich, L. and Gazit, E. (2008) Controlled patterning of peptide nanotubes and nanospheres using inkjet printing technology. Journal of Peptide Science, 14, 217±223. Albericio, F., Bofill, J. M., El-Faham, A. and Kates, S. A. (1998) Use of onium salt-based coupling reagents in peptide synthesis. Journal of Organic Chemistry, 63, 9678± 9683. Banwell, E. F., Abelardo, E. S., Adams, D. J., Birchall, M. A., Corrigan, A., Donald, A. M., Kirkland, M., Serpell, L. C., Butler, M. F. and Woolfson, D. N. (2009) Rational design and application of responsive -helical peptide hydrogels. Nature Materials, 8, 596±600. Birchall, L. S., Ulijn, R. V. and Webb, S. J. (2008) A combined SPS-LCD sensor for screening protease specificity. Chemical Communications, 2861±2863. Blackwell, H. E. (2006) Hitting the SPOT: small-molecule macroarrays advance combinatorial synthesis. Current Opinion in Chemical Biology, 10, 203±212. Boeneman, K., Mei, B. C., Dennis, A. M., Bao, G., Deschamps, J. R., Mattoussi, H. and Medintz, I. L. (2009) Sensing caspase 3 activity with quantum dotfluorescent protein assemblies. Journal of the American Chemical Society, 131, 3828±3829. Branco, M. C., Pochan, D. J., Wagner, N. J. and Schneider, J. P. (2009) Macromolecular diffusion and release from self-assembled beta-hairpin peptide hydrogels. Biomaterials, 30, 1339±1347. Carpino, L. A. (1993) 1-Hydroxy-7-azabenzotriazole. An efficient peptide coupling additive. Journal of the American Chemical Society, 115, 4397±4398. Carpino, L. A. and Han, G. Y. (1972) 9-Fluorenylmethoxycarbonyl amino-protecting group. The Journal of Organic Chemistry, 37, 3404±3409. Carpino, L. A., Shroff, H., Triolo, S. A., Mansour, E. M. E., Wenschuh, H. and Albericio, F. (1993) The 2,2,4,6,7-pentamethyldihydrobenzofuran-5-sulfonyl group (PBF) as arginine side-chain protectant. Tetrahedron Letters, 34, 7829±7832. Cho, Y. and Ivanisevic, A. (2004) SiOx surfaces with lithographic features composed of a TAT peptide. Journal of Physical Chemistry B, 108, 15223±15228. Cho, Y. and Ivanisevic, A. (2005) TAT peptide immobilization on gold surfaces: A comparison study with a thiolated peptide and alkylthiols using AFM, XPS, and FTIRRAS. Journal of Physical Chemistry B, 109, 6225±6232. Chow, D., Nunalee, M. L., Lim, D. W., Simnick, A. J. and Chilkoti, A. (2008) Peptidebased biopolymers in biomedicine and biotechnology. Materials Science and Engineering R-Reports, 62, 125±155. Coste, J., Lenguyen, D. and Castro, B. (1990) PYBOP ± a new peptide coupling reagent devoid of toxic by-product. Tetrahedron Letters, 31, 205±208. Dalsin, J. L., Hu, B. H., Lee, B. P. and Messersmith, P. B. (2003) Mussel adhesive protein mimetic polymers for the preparation of nonfouling surfaces. Journal of the American Chemical Society, 125, 4253±4258. Das, A. K., Collins, R. and Ulijn, R. V. (2008) Exploiting enzymatic (reversed) hydrolysis in directed setf-assembly of peptide nanostructures. Small, 4, 279±287. Dourtoglou, V., Ziegler, J. C. and Gross, B. (1978) Hexafluorophosphate of Obenzotriazole-N, N-tetramethyurea ± new and effective reagent of coupling peptide. Tetrahedron Letters, 1269±1272. Dreher, M. R., Simnick, A. J., Fischer, K., Smith, R. J., Patel, A., Schmidt, M. and Chilkoti, A. (2008) Temperature triggered self-assembly of polypeptides into multivalent spherical micelles. Journal of the American Chemical Society, 130, 687±694.

ß Woodhead Publishing Limited, 2011

Surface modification of biomaterials by peptide functionalisation

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El-Faham, A., Subiros, R. and Albericio, F. (2008) A non-explosive replacement for benzotriazole based coupling reagents. Journal of Peptide Science, 14, 57. Escuder, B. and Miravet, J. F. (2006) Silk-inspired low-molecular-weight organogelator. Langmuir, 22, 7793±7797. Estroff, L. A. and Hamilton, A. D. (2004) Water gelation by small organic molecules. Chemical Reviews, 104, 1201±1217. Falsey, J. R., Renil, M., Park, S., Li, S. J. and Lam, K. S. (2001) Peptide and small molecule microarray for high throughput cell adhesion and functional assays. Bioconjugate Chemistry, 12, 346±353. Fodor, S. P. A., Read, J. L., Pirrung, M. C., Stryer, L., Lu, A. T. and Solas, D. (1991) Light-directed, spatially addressable parallel chemical synthesis. Science, 251, 767± 773. Frank, R. (1992) Spot-synthesis ± an easy technique for the positionally addressable, parallel chemical synthesis on a membrane support. Tetrahedron, 48, 9217. Greene, T. W. and Wuts, P. G. M. (1999) Protective Groups in Organic Synthesis, New York, John Wiley & Sons Ltd. Haines-Butterick, L., Rajagopal, K., Branco, M., Salick, D., Rughani, R., Pilarz, M., Lamm, M. S., Pochan, D. J. and Schneider, J. P. (2007) Controlling hydrogelation kinetics by peptide design for three-dimensional encapsulation and injectable delivery of cells. Proceedings of the National Academy of Sciences, 104, 7791± 7796. Hartgerink, J. D., Beniash, E. and Stupp, S. I. (2002) Peptide-amphiphile nanofibers: A versatile scaffold for the preparation of self-assembling materials. Proceedings of the National Academy of Sciences of the United States of America, 99, 5133±5138. Heller, M. J. (2002) DNA microarray technology: devices, systems, and applications. Annual Review of Biomedical Engineering, 4, 129±153. Houseman, B. T., Gawalt, E. S. and Mrksich, M. (2003) Maleimide-functionalized selfassembled monolayers for the preparation of peptide and carbohydrate biochips. Langmuir, 19, 1522±1531. Izumiya, N. and Muraoka, M. (1969) Racemization test in peptide synthesis. Journal of the American Chemical Society, 91, 2391±2392. Jayawarna, V., Ali, M., Jowitt, T., Miller, A., Saiani, A., Gough, J. and Ulijn, R. (2006) Nanostructured hydrogels for three-dimensional cell culture through self-assembly of fluorenylmethoxycarbonyl-dipeptides. Advanced Materials, 18, 611±614. Jiang, H. Z. and Stupp, S. I. (2005) Dip-pen patterning and surface assembly of peptide amphiphiles. Langmuir, 21, 5242±5246. Karlstrom, A. and Unden, A. (1996) A new protecting group for aspartic acid that minimizes piperidine-catalyzed aspartimide formation in Fmoc solid phase peptide synthesis. Tetrahedron Letters, 37, 4243±4246. Kisiday, J., Jin, M., Kurz, B., Hung, H., Semino, C., Zhang, S. and Grodzinsky, A. J. (2002) Self-assembling peptide hydrogel fosters chondrocyte extracellular matrix production and cell division: Implications for cartilage tissue repair. Proceedings of the National Academy of Sciences of the United States of America, 99, 9996±10001. Kloxin, A. M., Kasko, A. M., Salinas, C. N. and Anseth, K. S. (2009) Photodegradable hydrogels for dynamic tuning of physical and chemical properties. Science, 324, 59±63. Laromaine, A., Koh, L. L., Murugesan, M., Ulijn, R. V. and Stevens, M. M. (2007) Protease-triggered dispersion of nanoparticle assemblies. Journal of the American Chemical Society, 129, 4156±4157. Laurent, N., Haddoub, R., Voglmeir, J., Wong, S. C. C., Gaskell, S. J. and Flitsch, S. L.

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(2008) SPOT synthesis of peptide arrays on self-assembled monolayers and their evaluation as enzyme substrates. Chembiochem, 9, 2592±2596. Liberski, A., Zhang, R. and Bradley, M. (2009) Inkjet fabrication of polymer microarrays and grids ± solving the evaporation problem. Chemical Communications, 334±336. Lutolf, M. P. and Hubbell, J. A. (2005) Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering. Nature Biotechnology, 23, 47±55. Mahler, A., Reches, M., Rechter, M., Cohen, S. and Gazit, E. (2006) Rigid, selfassembled hydrogel composed of a modified aromatic dipeptide. Advanced Materials, 18, 1365±1370. Mart, R. J., Osborne, R. D., Stevens, M. M. and Ulijn, R. V. (2006) Peptide-based stimuliresponsive biomaterials. Soft Matter, 2, 822±835. Mata, A., Hsu, L., Capito, R., Aparicio, C., Henrikson, K. and Stupp, S. I. (2009) Micropatterning of bioactive self-assembling gels. Soft Matter, 5, 1228±1236. Matysiak, S., Boldicke, T., Tegge, W. and Frank, R. (1998) Evaluation of monomethoxytrityl and dimethoxytrityl as orthogonal amino protecting groups in fmoc solid phase peptide synthesis. Tetrahedron Letters, 39, 1733±1734. McDonald, T. O., Qu, H. L., Saunders, B. R. and Ulijn, R. V. (2009) Branched peptide actuators for enzyme responsive hydrogel particles. Soft Matter, 5, 1728±1734. McKay, F. C. and Albertson, N. F. (1957) New amine-masking groups for peptide synthesis. Journal of the American Chemical Society, 79, 4686±4690. Medintz, I. L., Clapp, A. R., Brunel, F. M., Tiefenbrunn, T., Uyeda, H. T., Chang, E. L., Deschamps, J. R., Dawson, P. E. and Mattoussi, H. (2006) Proteolytic activity monitored by fluorescence resonance energy transfer through quantum-dot-peptide conjugates. Nature Materials, 5, 581±589. Meldal, M. (1992) PEGA ± a flow stable polyethylene-glycol dimethyl acrylamide copolymer for solid-phase synthesis. Tetrahedron Letters, 33, 3077±3080. Meldal, M. (2002) The one-bead two-compound assay for solid phase screening of combinatorial libraries. Biopolymers, 66, 93±100. Merrifield, R. B. (1963) Solid phase peptide synthesis. I. The synthesis of a tetrapeptide. Journal of the American Chemical Society, 85, 2149±2154. Montalbetti, C. and Falque, V. (2005) Amide bond formation and peptide coupling. Tetrahedron, 61, 10827±10852. Mosse, W. K. J., Koppens, M. L., Gengenbach, T. R., Scanlon, D. B., Gras, S. L. and Ducker, W. A. (2009) Peptides grafted from solids for the control of interfacial properties. Langmuir, 25, 1488±1494. Ozin, G. A. and Arsenault, A. C. (2005) Nanochemistry: A Chemical Approach to Nanomaterials, Cambridge, RSC Publishing. Palmer, L. C. and Stupp, S. I. (2008) Molecular self-assembly into one-dimensional nanostructures. Accounts of Chemical Research, 41, 1674±1684. Piehler, J., Brecht, A., Valiokas, R., Liedberg, B. and Gauglitz, G. (2000) A high-density poly(ethylene glycol) polymer brush for immobilization on glass-type surfaces. Biosensors and Bioelectronics, 15, 473±481. Piner, R. D., Zhu, J., Xu, F., Hong, S. H. and Mirkin, C. A. (1999) `Dip-pen' nanolithography. Science, 283, 661±663. Ramage, R. and Green, J. (1987) Ng-2,2,5,7,8-pentamethylchroman-6-sulfonyl-Larginine ± a new acid labile derivative for peptide-synthesis. Tetrahedron Letters, 28, 2287±2290. Rawsterne, R. E., Gough, J. E., Rutten, F. J. M., Pham, N. T., Poon, W. C. K., Flitsch, S. L., Maltman, B., Alexander, M. R. and Ulijn, R. V. (2006) Controlling protein

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retention on enzyme-responsive surfaces. Surface and Interface Analysis, 38, 1505± 1511. Rawsterne, R. E., Todd, S. J., Gough, J. E., Farrar, D., Rutten, F. J. M., Alexander, M. R. and Ulijn, R. V. (2007) Cell spreading correlates with calculated logP of amino acid-modified surfaces. Acta Biomaterialia, 3, 715±721. Reches, M. and Gazit, E. (2006) Controlled patterning of aligned self-assembled peptide nanotubes. Nature Nanotechnology, 1, 195±200. Reineke, U., Volkmer-Engert, R. and Schneider-Mergener, J. (2001) Applications of peptide arrays prepared by the SPOT-technology. Current Opinion in Biotechnology, 12, 59±64. Reynolds, N. P., Tucker, J. D., Davison, P. A., Timney, J. A., Hunter, C. N. and Leggett, G. J. (2009) Site-specific immobilization and micrometer and nanometer scale photopatterning of yellow fluorescent protein on glass surfaces. Journal of the American Chemical Society, 131, 896±897. Salaita, K., Wang, Y. H. and Mirkin, C. A. (2007) Applications of dip-pen nanolithography. Nature Nanotechnology, 2, 145±155. Salisbury, C. M., Maly, D. J. and Ellman, J. A. (2002) Peptide microarrays for the determination of protease substrate specificity. Journal of the American Chemical Society, 124, 14868±14870. Sheehan, J. C. and Hess, G. P. (1955) A new method of forming peptide bonds. Journal of the American Chemical Society, 77, 1067±1068. Sheehan, J. C., Goodman, M. and Hess, G. P. (1956) Peptide derivatives containing hydroxyamino acids. Journal of the American Chemical Society, 78, 1367±1369. Sheehan, J., Cruickshank, P. and Boshart, G. (1961) Notes ± a convenient synthesis of water-soluble carbodiimides. The Journal of Organic Chemistry, 26, 2525±2528. Sieber, P. and Riniker, B. (1991) Protection of carboxamide functions by the trityl residue ± application to peptide-synthesis. Tetrahedron Letters, 32, 739±742. Silva, G. A., Czeisler, C., Niece, K. L., Beniash, E., Harrington, D. A., Kessler, J. A. and Stupp, S. I. (2004) Selective differentiation of neural progenitor cells by highepitope density nanofibers. Science, 303, 1352±1355. Smith, A. M., Williams, R. J., Tang, C., Coppo, P., Collins, R. F., Turner, M. L., Saiani, A. and Ulijn, R. V. (2008) Fmoc-diphenylalanine self assembles to a hydrogel via a novel architecture based on pi-pi interlocked beta-sheets. Advanced Materials, 20, 37±41. Stevens, M. M. and George, J. H. (2005) Exploring and engineering the cell surface interface. Science, 310, 1135±1138. Thornton, P. D., Mart, R. J., Webb, S. J. and Ulijn, R. V. (2008) Enzyme-responsive hydrogel particles for the controlled release of proteins: designing peptide actuators to match payload. Soft Matter, 4, 821±827. Todd, S. J., Farrar, D., Gough, J. E. and Ulijn, R. V. (2007) Enzyme-triggered cell attachment to hydrogel surfaces. Soft Matter, 3, 547±550. Todd, S. J., Scurr, D. J., Gough, J. E., Alexander, M. R. and Ulijn, R. V. (2009) Enzymeactivated RGD ligands on functionalized poly(ethylene glycol) monolayers: surface analysis and cellular response. Langmuir, 25, 7533±7539. Toledano, S., Williams, R., Jayawarna, V. and Ulijn, R. (2006) Enzyme-triggered selfassembly of peptide hydrogels via reversed hydrolysis. Journal of the American Chemical Society, 128, 1070±1071. Ulijn, R. V. and Smith, A. M. (2008) Designing peptide based nanomaterials. Chemical Society Reviews, 37, 664±675. Ulijn, R. V., Baragana, B., Halling, P. J. and Flitsch, S. L. (2002) Protease-catalyzed

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peptide synthesis on solid support. Journal of the American Chemical Society, 124, 10988±10989. Valeur, E. and Bradley, M. (2009) Amide bond formation:beyond the myth of coupling reagents. Chemical Society Reviews, 38, 606±631. Wang, Z., Levy, R., Fernig, D. G. and Brust, M. (2005) The peptide route to multifunctional gold nanoparticles. Bioconjugate Chemistry, 16, 497±500. Wegner, G. J., Lee, H. J. and Corn, R. M. (2002) Characterization and optimization of peptide arrays for the study of epitope-antibody interactions using surface plasmon resonance imaging. Analytical Chemistry, 74, 5161±5168. Welser, K., Grilj, J., Vauthey, E., Aylott, J. W. and Chan, W. C. (2009) Protease responsive nanoprobes with tethered fluorogenic peptidyl 3-arylcoumarin substrates. Chemical Communications, 671±673. Williams, R. J., Smith, A. M., Collins, R., Hodson, N., Das, A. K. and Ulijn, R. V. (2009) Enzyme-assisted self-assembly under thermodynamic control. Nature Nanotechnology, 4, 19±24. Wilson, D. L., Martin, R., Hong, S., Cronin-Golomb, M., Mirkin, C. A. and Kaplan, D. L. (2001) Surface organization and nanopatterning of collagen by dip-pen nanolithography. Proceedings of the National Academy of Sciences of the United States of America, 98, 13660±13664. Windridge, G. and Jorgensen, E. C. (1971) 1-Hydroxybenzotriazole as a racemizationsuppressing reagent for incorporation of IM-benzyl-L-histidine into peptides. Journal of the American Chemical Society, 93, 6318±6319. Woolfson, D. N. and Ryadnov, M. G. (2006) Peptide-based fibrous biomaterials: some things old, new and borrowed. Current Opinion in Chemical Biology, 10, 559±567. Yang, Z. M., Gu, H. W., Fu, D. G., Gao, P., Lam, J. K. and Xu, B. (2004) Enzymatic formation of supramolecular hydrogels. Advanced Materials, 16, 1440±1444. Yang, Z., Liang, G. and Xu, B. (2008) Enzymatic hydrogelation of small molecules. Accounts of Chemical Research, 41, 315±326. Yao, H. Q., Zhang, Y., Xiao, F., Xia, Z. Y. and Rao, J. H. (2007) Quantum dot/ bioluminescence resonance energy transfer based highly sensitive detection of proteases. Angewandte Chemie-International Edition, 46, 4346±4349. Yokoi, H., Kinoshita, T. and Zhang, S. G. (2005) Dynamic reassembly of peptide RADA16 nanofiber scaffold. Proceedings of the National Academy of Sciences of the United States of America, 102, 8414±8419. Zhang, R., Liberski, A., Khan, F., Diaz-Mochon, J. J. and Bradley, M. (2008) Inkjet fabrication of hydrogel microarrays using in situ nanolitre-scale polymerisation. Chemical Communications, 1317±1319. Zhang, S., Holmes, T., Lockshin, C. and Rich, A. (1993) Spontaneous assembly of a selfcomplementary oligopeptide to form a stable macroscopic membrane. Proceedings of the National Academy of Sciences of the United States of America, 90, 3334± 3338. Zhang, S. G. (2003) Fabrication of novel biomaterials through molecular self-assembly. Nature Biotechnology, 21, 1171±1178. Zhang, S. G., Yan, L., Altman, M., Lassle, M., Nugent, H., Frankel, F., Lauffenburger, D. A., Whitesides, G. M. and Rich, A. (1999) Biological surface engineering: a simple system for cell pattern formation. Biomaterials, 20, 1213±1220. Zhang, Y., Gu, H., Yang, Z. and Xu (2003) Supramolecular hydrogels respond to ligandreceptor interaction. Journal of the American Chemical Society, 125, 13680±13681. Zhao, X. J. and Zhang, S. G. (2006) Molecular designer self-assembling peptides. Chemical Society Reviews, 35, 1105±1110.

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Zhou, M., Smith, A. M., Das, A. K., Hodson, N. W., Collins, R. F., Ulijn, R. V. and Gough, J. E. (2009) Self-assembled peptide-based hydrogels as scaffolds for anchorage-dependent cells. Biomaterials, 30, 2523±2530. Zourob, M., Gough, J. E. and Ulijn, R. V. (2006) A micropatterned hydrogel platform for chemical synthesis and biological analysis. Advanced Materials, 18, 655±659.

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Metal surface oxidation and surface interactions L . D E N A R D O , G . R A F F A I N I , F . G A N A Z Z O L I and R . C H I E S A , Politecnico di Milano, Italy

Abstract: In this chapter we discuss the main aspects of metal oxides in metallic biomaterials, with a special emphasis on titanium. Basic chemicophysical principles, in terms of titanium oxide layer ability to promote specific interactions in biological environments, are discussed. Further on, the theoretical approaches to describe the surface behaviour at atomistic level are reviewed, together with the hydration studies and protein adsorption at the surface of the titanium dioxide polymorphs. We then critically discuss the role of chemical and electrochemical surface modification technologies, aimed at modifying TiO2 structure, morphology and chemistry to tailor in vivo biological response. A perspective analysis of the correlation between theoretical studies and metal oxide surface modifications is finally offered, discussing their strengths and weaknesses and emphasizing their role as a powerful design tool for a new generation of implantable devices in which metal oxide surfaces can be tuned to yield a specific biological response. Key words: biomaterial surface modifications, theoretical modelling, metal oxide hydration, surface adsorption, implantable devices.

5.1

Surface oxides in metallic medical devices: the scenario

Metals and metal alloys find widespread application in medical devices, since they are reliable in terms of the stringent mechanical performances required in biological function substitution and/or support. This class of materials is extensively used in medicine, especially for load bearing and hard tissue implantable prostheses. In Fig. 5.1 some metal orthopaedic and dental devices are shown, including knee, hip and dental prostheses (photo courtesy of SAMO Biomedica SpA). Historically, metal alloys have been successfully and systematically introduced during the twentieth century and mainly used in orthopaedic applications (Navarro et al., 2008): stainless steel and Co-based alloys have been applied in hard (bone plates, total and partial joint replacement, dental implants, etc.) and soft tissue substitution/support (metal stents, components of artificial hearts, pacemakers, etc.). After this first generation of metal implants, an increasing number of devices has been realized in titanium and titanium alloys (Long and Rack, 1998).

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5.1 Metal orthopaedic and dental prosthetic components (courtesy of SAMO S.p.A., Italy).

In Table 5.1 the main metal alloys for biomedical applications are reported. Describing physicochemical and mechanical properties of metal alloys is beyond the main scope of this chapter, and we suggest the reader to refer to the several excellent published books, chapters and review articles about this subject (see, for instance, the works of Brunski, 1996; Freese et al., 2001 and Niinomi, 2008). The historical perspective of metal alloy applications in medical devices can be seen as a paradigm of biomaterial evolution throughout different generations: the availability of an appropriate material is correlated to the advancements in materials science and technology, becoming a challenging issue for material scientists. Along this direction a pre-eminent role is played by the metal oxides that characterize the surface of biomedical metal alloys. The metal implant response is, in fact, mediated by an oxide layer that covers the metal surface and depends on the bulk chemical composition. Four main oxides can be found on biomedical metal alloys, namely: (i) Cr2O3, chromium oxide on stainless steels and cobalt chromium alloys, (ii) TiO2, titanium dioxide on titanium and its alloys, (iii) Nb2O5, niobium oxide on niobium, and (iv) Ta2O5, tantalum oxide on tantalum. In the case of alloys, the natural oxide film can be also composed by some oxides of the alloying elements: for instance, Al2O3 is reported in Ti6Al4V alloys (Textor et al., 2001). Vanadium has been also reported in +III, +IV, and +V oxidized forms in the outer oxide layer. The compact oxide layer on biomedical metal alloys plays a primary role in insulating the reactive underlying metal from the external environment, preventing further corrosion.

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Table 5.1 Some characteristics of orthopaedic metallic implant materials (modified from Long and Rack, 1998) Material Class

Elements Designation

Weight (%)

Mechanical properties E (GPa)

Stainless F-138 Cr (17±20) 205±210 steel (316 LDVM) Ni (12±14) Mo (2±4) Fe (bal.)

Y (MPa)

UTS (MPa)

170±750

465±950

Use

Temporary devices Stents

Co alloys F-75 F-799 F-1537

Cr (19±30) 220±230 275±1585 Mo (0±10) Ni (0±37) Co (bal.)

600±1785 Implantable prostheses Articulating surfaces

Ti alloys

Al (0±6) V (0±4) Nb (0±40) Ta (0±5) O (0±0.5) Zr (0±13) Ni (0±55) Ti (bal.)

590±1024 Implantable prostheses Long-term permanent devices

F-67 F-136 F-1295

20±110

50±921

F Designations are according to ASTM

All currently used alloys are subjected to corrosion when in contact with body fluids as the body environment is very aggressive: many forms of corrosion could take place depending on the environmental composition, pH and metal nature, namely pitting, crevice, fretting corrosion, etc. (Geetha et al., 2009). Corrosion processes of metals at the implant site are determined by both thermodynamic driving forces, which cause corrosion (oxidation and reduction) reactions, and kinetic barriers, which limit the rate of these reactions. Although there are many forms of corrosion damage, the rate of attack of general corrosion is very low: in a physiological environment stainless steel, Coand Ti-based alloys show corrosion rates that are extremely low and allow them to be used as the three main classes of implantable biomedical devices. This corrosion resistance is due to the formation of a passive film, which is influenced by a combination of the various constituents of the environment in which the device is manufactured and subsequently operates: it represents the main kinetic barrier to corrosion and impedes or prevents corrosion reactions from taking place (Balamurugan et al., 2008). Moreover, properties of this layer can be further modulated in order to improve corrosion resistance, via some of the technologies further presented in this chapter. Cigada et al. (1992), for instance, demonstrated it via anodization of Ti alloys.

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Metal alloys found extensive use in partial or total joint replacement, in which requirements for mechanical stability are coupled with tribological performance. Since natural joints have remarkable tribological characteristics, devices replacing these functions have to be designed in order to minimize wear by optimizing lubrication and compliance. The surface features of metal alloys are, in fact, of primary importance for friction and wear resistance and in lubrication. Several clinical studies have shown that excessive wear of joint replacing materials is the principal cause of failure in joint prostheses and a large amount of research has been conducted with the aim of studying the wear mechanisms for two surfaces moving relatively (Blunt et al., 2009): the involved mechanisms were explained, looking at abrasive and adhesive wear. A systematic study (McGee et al., 2000) of the extent of wear in total hip replacement with CoCrMo, stainless steel or Ti6Al4V femoral heads articulated against Ultra High Molecular Weight Polyethylene (UHMWPE) showed a significant difference among these materials. Titanium alloy femoral heads had the higher wear and CoCrMo heads had the lower wear. Consistent with these data, titanium articulating surfaces are no longer recommended for clinical use. Similarly to corrosion resistance, the importance of the oxide layer in articulating device responses is confirmed by the attempt to enhance the wear resistance of some materials via oxidation processes: diffusion/oxidation surface hardening has been shown to be an interesting technique along this direction (Streicher et al., 1991). With the introduction of titanium alloys it became evident that surface oxides not only play a primary role in corrosion and wear resistance, but they are able to modulate specific interaction response in vivo with the contact biological tissues. Even if historically they have been exploited because of their lower modulus, superior biocompatibility and enhanced corrosion resistance (Long and Rack, 1998), when interacting with biological fluids and their constituents, the surfaces of titanium and titanium alloys show distinctive features. Understanding the specificity of TiO2 in biological fluid interactions has then become a primary goal in biomaterial interface science: theoretical and computational studies could help to explain the fundamentals of these processes, leading the emerging technologies for surface modification to tailor a new generation of medical devices with advanced functionalities.

5.2

Titanium oxides on Ti implants: from crystallographic structure to the theoretical study of the atomistic surface structure and behaviour

As pointed out in the previous section, metal oxides that naturally cover the surface of metal alloys commonly used in biomedical devices play a primary role in modulating corrosion and wear resistance. The ability of being protected

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by a passivating oxide film had been the original motivation for stainless steel and CoCrMo alloys to be applied in implantable devices, especially for loadbearing applications (Brunski, 1996). Since the first successful applications of titanium in modern medicine during 1960s (Pilliar, 1987, 1998), Ti alloys have gradually replaced other metals in applications with direct bone contact and high stress distribution, such as orthopaedic and dental implants. In fact, Ti alloys fulfil, for bone-contact implants, biomaterial design requirements in a more efficient way than other metal alloys as a consequence of a good combination of mechanical and biocompatibility properties, together with in vivo chemical and biological stability (Liu et al., 2004). One of the most important parameters of this success is correlated to the layer of titanium oxide (mainly TiO2) which spontaneously forms on the Ti surface. Understanding the composition, the microstructure and the physico-chemical properties of titanium oxides is a central point to explain the interactions with the biological environment: several studies have been devoted to this subject, mainly because of the wide technological interest in titanium oxides due to their catalytic, electronic and pigment properties (Diebold, 2003). In this section, after a preliminary description of the chemico-physical properties of titanium oxide, we will focus on the state-of-art description of computational studies devoted to understanding the surface structure of TiO2 and its response when in contact with water or a biological environment, with special emphasis on crystalline TiO2 polymorphs. Section 5.3 will be devoted to the main technologies that allow structuring of the titanium alloy surfaces to improve the overall biological response, according to the input deriving from computational studies.

5.2.1

Titanium oxides on titanium alloys

A thin titanium oxide layer naturally covers titanium alloys and it is generally present in its more thermodynamically stable form, i.e. TiO2 (oxidation state +IV). However, titanium oxide is a class of oxides with a wide range of Ti/O ratios, namely Ti3O, Ti3O2, TiO, Ti2O3, Ti3O5, TiO2, Ti2O2 (Textor et al., 2001): the presence of several oxides on titanium surfaces depends both on the different titanium oxidation states and the oxygen solubility in Ti, that leads to a gradual change of this ratio from the outer part of the surface (+IV oxidation state) to the bulk of the material (metal). Moreover, depending on the alloying materials, several oxides have been detected on titanium alloy surfaces (Long and Rack, 1998): · Al2O3 on Ti6Al4V · Al2O3 and Nb2O5 on Ti6Al7Nb However, TiO2 is the predominant oxide on titanium alloys, and its properties modulate the biological interactions when a Ti-based device is implanted in human body. TiO2 found widespread technological application because it shows

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interesting properties: it is a non-stoichiometric n-type semiconductor oxide used for application ranging from heterogeneous photo-induced catalysis to antifogging surfaces. Chemico-physical properties of TiO2 have been described in several excellent reviews: readers can refer to the excellent and monumental reviews of Diebold (2003) and Carp et al. (2004).

5.2.2

Theoretical description of TiO2 surfaces: quantum mechanical studies of the crystallographic surface in vacuo

In this chapter we are interested in understanding the behaviour of TiO2 as passive metal oxide film naturally covering in a physiological environment the surface of titanium used as a biomaterial. Nevertheless we think it could be useful first to describe the structure of titanium oxide placed in vacuo or in air. A significant amount of theoretical work was carried out to describe the structure and possible reconstruction of some low index surfaces of TiO2 mostly in the rutile phase, but also in some cases in the anatase polymorph phase, usually considering the surfaces most thoroughly characterized experimentally (Diebold, 2003; Henderson, 2002). The TiO2 surfaces are of paramount interest for their photocatalytic activity in air, and accordingly most studies dealt with the atomistic description of the surfaces in vacuo, exposing in general bridging 2-fold coordinated oxygen atoms and 5-fold coordinated titanium atoms showing a vacancy when compared to the bulk 6-fold coordinated atoms. This is the case of the most stable TiO2 surface, rutile (1 1 0) shown in Fig. 5.2. Accordingly, instead of the formal Ti4+ charge of the bulk titanium oxide, the exposed titanium has a formal Ti3+ charge, related to its catalytic activity. Indeed, recent theoretical calculations for rutile (1 1 0) (Labat et al., 2008) indicate a charge of +2.70, calculated by the simple Hartree-Fock method, or of +2.31, calculated by the more accurate Density Functional Theory (DFT). For anatase (1 0 1) a charge of +2.65 (Hartree-Fock) or +2.27 (DFT) was calculated by the Mulliken charge analysis. Conversely, in water or in physiological fluids,

5.2 The most stable TiO2 surface, rutile (1 1 0). Oxygen and titanium atoms are shown in dark and light grey spheres, respectively.

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the exposed TiO2 surfaces may show hydroxyl groups at the bridging oxygen atoms (denoted also as BOH groups in the following) or as terminal groups on the 5-fold coordinated Ti atoms (denoted as TOH groups, or as protonated TOH2 for an undissociated water molecule). We shall discuss this issue in the next paragraph, where the structure of the TiO2 surfaces in water, together with their hydration in the bulk liquid, and the presence of the surface hydroxyls will be examined in more detail. The surface theoretically modelled in vacuo by many authors was mostly low index rutile (1 1 0), because it is the most stable one among the TiO2 polymorphs with the lowest surface energy, and because of its detailed experimental characterization with a variety of techniques. Other surfaces theoretically studied in detail for rutile were also (1 0 0), showing the largest surface energy (Diebold, 2003) and (0 0 1). The freshly cleaved rutile (1 1 0) undergoes some reconstruction compared to the ideal bulk geometry, related with the presence of exposed bridging 2-fold coordinated oxygen atoms and 5-fold coordinated titanium atom, in addition to 3-fold coordinated oxygen atoms and 6-fold coordinated titanium atoms that characterize the bulk metal oxide. Early surface X-ray diffraction (SXRD) experiments (Charlton et al., 1997) indicated quite large displacements of the bridging oxygen atoms and of the exposed titanium atoms towards the bulk, unlike the 3-fold coordinated surface oxygen atoms, for instance, due to the surface reconstruction minimizing the surface energy. Theoretical calculations of this surface were carried out by many authors (see, for instance, Harrison et al. (1999) and more recently Swamy et al. (2002) and Labat et al. (2008)). These works were in reasonable agreement among themselves, but yielded a result quite different from the experimental SXRD data, suggesting significantly smaller surface deformations mostly in the opposite direction, even though by a very small amount, for the bridging oxygen atoms when compared to the experimental results. Interestingly, such theoretical predictions were consistently obtained both through Hartree-Fock methods using an LCAO approximation (linear combination of atomic orbitals), and through the more accurate DFT technique to directly calculate the electron density. The simulated systems comprised in all cases thin slabs having the thickness of only a few Ti-O layers, which could quantitatively somewhat affect the results, but not the general qualitative features. In particular, Swamy et al. (2002) reported an extensive computational study of rutile (1 1 0) using both Hartree-Fock and DFT methods with periodic boundary conditions parallel to the surface directions, and checked for the effect of the slab thickness, and for a possible influence of the chosen basis set (i.e., the shape and number of atomic orbitals of the Ti and O atoms used in the calculations), but the general disagreement with the SXRD experiments of Charlton et al. (1997) did remain. Much effort was spent to find the source of this discrepancy, considering, for instance, finite temperature effects of the experimental data, or the poor estimate of the (fractional) ionic charges of the surface atoms in the theoretical methods, or else

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a low frequency phonon mode associated with the bridging oxygen atoms, leading to their displacement even at low temperatures (Harrison et al., 1999). In the last few years, however, new experimental results have shed light on this issue, fully vindicating the theoretical results. In fact, low energy electron diffraction (LEED) experiments (Lindsay et al., 2005) unambiguously showed the correctness of the theoretical results obtained in particular by Swamy et al. (2002) and the minor relevance of the postulated low frequency phonons, at least at a low temperature. As a consequence, a new experimental determination of the structure of rutile (1 1 0) was carried out by the SXRD technique on a carefully prepared sample (Cabailh et al., 2007). A thorough analysis of the observed data unambiguously showed that the bridging oxygen atoms displayed an outward displacement with respect to the bulk material, unlike the previous SXRD result, but in good agreement with theoretical calculations. The disagreement with the previous SXRD data was attributed to a poorer quality of the surface and to possible experimental artefacts. It should also be noted that the LEED experiments used a very low incident electron beam current to avoid surface damage during data acquisition, a fact suggesting that this could also be an experimental problem also with other techniques possibly leading to surface reconstruction. The huge interest in rutile (1 1 0) produced a large number of papers devoted to its experimental and theoretical study, and here we only summarize some recent significant papers, while an extensive discussion of previous studies can be found in many recent reviews (Carp et al., 2004; Diebold, 2003; Thompson and Yates, 2006). As anticipated, the interest in this surface arises from the fact that it is the most stable rutile surface, and the presence of surface `vacancies' due to the 2-fold coordinated bridging oxygen atoms and the 5-fold coordinated titanium affecting its electronic properties and catalytic properties (Thompson and Yates, 2006). In particular, these vacancies may catalytically drive the dissociation of many molecules, in particular of water even in high vacuum, possibly giving rise to surface hydroxyl groups that are important for a fully hydrated material, in particular on the TiO2 surfaces of titanium biomaterials in a physiological environment, as described in the next paragraph. Other bare TiO2 surfaces were theoretically studied by Hartree-Fock and DFT methods considering other low-index rutile surfaces such as (1 0 0) (Lindan et al., 1996) and (0 0 1) having a relatively low stability and tending to reconstruct (Muscat and Harrison, 2000). More recently, the structure and the electronic properties of these surfaces were theoretically investigated anew (Labat et al., 2008) for a consistent check of the results with the same methodology, and for a more significant comparison with additional results for selected low-index anatase surfaces, namely (1 0 1) and (1 0 0). These calculations were carried out with high-level Hartree-Fock and DFT methods using a large basis set and two different computational approaches for the latter methodology, considering also the influence of the thickness of the considered

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slab. The geometrical results (i.e. the displacement of the exposed atoms with respect to the positions of a freshly cleaved surface) were in good agreement with previous theoretical and experimental results, whenever available. Most interestingly, however, the authors did compare the surface energy of the different rutile and anatase surfaces, determining their relative stability. Thus, while rutile (0 0 1) was shown to be the least stable among the five surfaces analyzed in the paper, rutile (1 1 0) and anatase (1 0 1) turned out to be the most stable ones, but their relative stability was somewhat dependent on the adopted computational methodology. As a conclusion of this section on non-hydroxylated TiO2 surfaces, we also mention a recent investigation of rutile (0 1 1) (Kubo et al., 2007). Both noncontact AFM and STM experiments showed this rutile surface to be composed of (2  1) and (4  1) structures, and the theoretical calculations were mainly used to calculate the stability and the electron density of these reconstructed surfaces.

5.2.3

Hydration phenomena on TiO2 surfaces: quantum mechanical and molecular dynamics (MD) simulation studies of water adsorption

The catalytic interaction of water molecules at the TiO2 surface was recently investigated theoretically. This situation does not correspond to a fully hydrated surface immersed in water or in a physiological fluid, but still it provides a way to study the possible formation of surface hydroxyls. In fact, as clearly pointed out by Machesky et al. (2008) in a comprehensive theoretical and experimental study of rutile (1 1 0) in contact with bulk liquid water, `the under-coordinated cation and oxygen atoms at most oxide surfaces are reactive toward water. In broad terms this reactivity can be classified as dissociative if first-layer water dissociates, or associative if it does not (within the time frame and P, T conditions of the observation). One common result of this reactivity toward water is that surface relaxation and reconstruction effects are often less extensive on hydrated than corresponding vacuum-terminated oxide surfaces. Another is surface protonation whereby hydrated metal oxide surfaces adsorb or desorb protons and hence acquire charge in response to bulk solution pH' (Machesky et al., 2008). Unfortunately, detecting and/or predicting the presence of physisorbed water molecules, the distribution of bridging and terminal hydroxyls (BOH and TOH, or TOH2, respectively) and their possible loss of an acidic proton producing a surface with a non-vanishing charge density is not trivial due to the small electron density on hydrogen that hampers its detection with X-rays, for instance. Therefore, theoretical modelling of the surface interaction with water may provide important information, or at least relevant clues for understanding the surface hydration. On the other hand, the surface hydration of TiO2 can have

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relevant consequences, in addition to affecting the performance of titanium as a biomaterial. For instance, hydration affects the stability of anatase and its conversion to rutile: while the former polymorph is more stable than the latter at the nanoscale, it converts to rutile when the nanocrystal exceeds a size of about 14 nm in vacuo, but at a somewhat larger size in water, even though the transition is also affected by impurities, the reaction atmosphere and the synthesis conditions (see, for instance, Barnard et al. (2005) and references therein). Also, UV irradiation strongly affects the hydrophilicity of TiO2, with important consequences for practical applications such as self-cleaning surfaces. In fact, Wang et al. (1997) reported a reversible UV-induced hydrophilicity on TiO2 for polycrystalline or single-crystal samples of both rutile and anatase. Such changes in the wettability behaviour were originally attributed to a radiation-induced loss of bridging surface oxygen atoms, producing a substantial amount of exposed Ti atoms (formally Ti3+), which would induce dissociative adsorption of water. More recently, however, this induced hydrophilicity effect was shown to be due to the effect of hydrocarbon contaminants in the atmosphere that were removed upon UV irradiation, but slowly redeposited in the dark (Thompson and Yates, 2006). A large body of work was carried out for the interaction of water vapour with TiO2 surfaces, extensively reviewed in the last decade or so by Henderson (2002) and by Diebold (2003), and therefore here we mostly focus on more recent theoretical calculations. These works mostly dealt with rutile (1 1 0), but some papers also reported results for rutile (1 0 0) and anatase (1 0 1) and (0 0 1). A central issue, which still appears to be open to discussion to some extent, is whether water adsorption on these surfaces is dissociative or molecular. While water adsorption involves both processes on rutile (1 0 0) as experimentally found and theoretically predicted, there is disagreement for rutile (1 1 0). In fact, in this case most calculations predicted significant water dissociation, whereas the experiments mostly suggested an unlike behaviour, with molecular adsorption only, except at defect sites (Diebold, 2003). A rather comprehensive study of cation adsorption on rutile (1 1 0) in bulk water was recently reported (Zhang et al., 2004). In this paper synchrotron X-ray reflectivity and X-ray standing wave measurements together with ion adsorption data from rutile powders were combined with different computational studies. The latter results were first obtained by DFT calculations of the surface with Sr2+ and Zn2+ ions respectively surrounded by three and five water molecules in their first hydration shell, while the other coordination sites of the cations were occupied by the surface oxygen atoms. The distribution of the water molecules and of the ions was then analyzed through classical MD simulations in the presence of a bulk solution. In water, the bridging oxygen atoms are expected to be either unprotonated or singly protonated (BO and BOH), the three-fold coordinated oxygen atoms in the Ti-O surface plane to be unprotonated, and the terminal oxygen atoms over the five-fold coordinated Ti atoms to be either

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singly or doubly protonated (TOH and TOH2, respectively). In the theoretical study, the partially hydrated cations were placed in vacuo close to the rutile (1 1 0) surface of a thin slab consisting of three layers, the central one being fixed, and the other ones being free to adjust. Minimum energy structures were determined by DFT, and the geometric and energetic parameters were then used in the force field adopted for the MD simulations, considering a four-layer TiO2 slab surrounded by the bulk solution. These simulations included either an optimized surface exposing only bridging oxygens and five-fold coordinated Ti atoms as in vacuo, or fully protonated at the bridging and terminal hydroxyls (i.e., BOH and TOH), yielding an electrically neutral surface. Furthermore, a slightly negative surface charge density was generated by removing some protons to the water molecules trapped near the exposed Ti atoms. In fact, we recall that the classical MD simulations cannot account for possible chemical reaction, involving, for instance, the surface dissociation of water or the hydroxyl protonation or deprotonation. The X-ray reflectivity data indicated that the surface presented bridging and terminal oxygen atoms with minor distortions with respect to the ideally cleaved bulk structure. A layer of adsorbed water Ê above the surface atoms; however, since molecules was detected at about 3.8 A X rays can hardly detect the hydrogen atoms, the protonation state of the surface oxygen atoms could not be determined. On the other hand, the MD simulations allowed estimation of the water density near the surface, in particular the water molecules tightly coordinated to the exposed Ti atoms, and slightly displaced at larger distances on the negatively charged surface. Additionally, a broad Ê from the Ti±O plane due maximum of the water density was found at 3.5±4.0 A to the first hydration shell (Zhang et al., 2004). A more thorough theoretical study of the hydration and protonation of rutile (1 1 0) was recently carried out with a combination of DFT, classical and quantum MD (Machesky et al., 2008). The quantum DFT method was used to optimize the surface and its interaction with a few water molecules, providing a static picture of the surface that was subsequently adopted as input for MD simulations. The classical MD simulations were performed after introducing the bulk water on the surfaces of a TiO2 slab, and the distribution of water molecules was determined in this way. The first hydration layer (one molecule thick) was then used as input for the quantum MD simulations, where forces were obtained from DFT calculations. The latter method, however, gave only a quasi-static picture of the hydration because the relaxation times of the water molecules exceeded the length of the accessible simulation time due to the intermolecular hydrogen bonds. The main practical difference between classical and quantum MD is that the latter procedure can only deal with a very limited surface hydration unlike the former, which allows inclusion of a large number of water molecules that cannot, however, react and/or dissociate. Accordingly, the classical MD runs were carried out for both hydroxylated and non-hydroxylated surfaces similar to the bare surface exposed in vacuo (Machesky et al., 2008).

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The analysis of the oxygen-hydrogen pair correlation function, yielding the (relative) probability density as a function of the H


O separation, provided the statistical distribution of these distances sampled in the dynamic trajectories, showing that the hydrogen bonds of water with the TOH and BOH hydroxyls were stronger than those between two water molecules. In particular, water best interacts with terminal hydroxyls at room temperature, both in terms of their average number per surface site (about 1.5 on TOH and 0.8 on BOH) and in terms of their strength. Moreover, the H-bonds on the TOH mostly involved a surface oxygen and a water hydrogen, while the opposite pattern was found for BOH. The hydration of other TiO2 surfaces, including also the surfaces of the various polymorphs, has been little studied. One notable exception is a comparative study by quantum DFT methods (Barnard et al., 2005) of the anatase (0 0 1), (1 0 0) and (1 0 1) and rutile (1 0 0), (0 1 1) and (1 1 0), the latter being considered for a comparison with previous work discussed before. One key point is that rutile is thermodynamically the most stable polymorph, at ambient conditions, but anatase is more commonly found in nanoscale samples, where it is more stable up to the size of about 14 nm in vacuo, and it converts to rutile only at very large temperatures. Therefore, it is of great interest to understand the structural stability of the two polymorph nanocrystals in water, which can be obtained by their surface energies and the minimum energy morphology. The surface hydration was investigated (Barnard et al., 2005) terminating a cleaved surface of the bulk material with a monolayer of water molecules so as to produce an electrically neutral surface by adsorption of a whole or of a dissociated water molecule. In this way, no real wetting of the surface was actually considered, and therefore the possible presence of TOH and BOH hydroxyls was determined by the calculated water behaviour. On the anatase (0 0 1) and (1 0 0) surfaces the adsorption was dissociative, producing the protonation of bridging oxygen atoms and addition of a terminal OH to the exposed 5-fold coordinated Ti atoms, whereas the anatase (1 0 1) surface showed only adsorption of a whole water molecule on Ti. An interesting issue is that the surface reconstruction of the hydrated surface is much smaller than in vacuo, so that hydration leads to minor displacements of the surface atoms with respect to the ideal geometry obtained by simply cleaving the bulk TiO2. The rutile surfaces were also predicted to show dissociative adsorption of water monolayer on the (1 0 0) and the (0 1 1) surface, and molecular adsorption on the (1 1 0) surface. In the two former cases, dissociative adsorption led to OH-terminated Ti atoms and protonation of the bridging oxygen atoms, whereas in the latter case the water adsorption completed the coordination of the exposed Ti atoms. Unlike what was found for the anatase surfaces, in rutile the surface reconstruction due to the first hydration layer was in any case predicted to be usually slightly larger than in vacuo, in particular for the (0 1 1) surface. The adsorption energy obtained by these calculations was further applied to get the surface free

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energy and surface tension for all the above-mentioned cases. The relative stability of the various surfaces turned out to be the same in water and in vacuo, but with significant differences in the individual values, while the surface tension showed much larger differences. In any case, based on the calculated values and considering the expected crystal morphologies, nanosized anatase was predicted to be thermodynamically more stable than rutile both in vacuo and in water, and the limiting size for thermodynamically stable anatase nanocrystals was expected to be larger in water than in vacuo. These calculations considered the interactions and possible reactions of some rutile and anatase surfaces in terms of the chemisorption and reactivity of a few water molecules. On the other hand, simulations of a fully immersed TiO2 surface, are mostly lacking in the literature, with the partial exception of the rutile (1 1 0) surface discussed before (Machesky et al., 2008). Very recently, however, the issue of physisorption of larger molecules, including mono- and dipeptides (Carravetta and Monti, 2006; Monti et al., 2008) and lipids (Fortunelli and Monti, 2008) on TiO2 have begun to be addressed by classical MD simulations. These studies considered the adsorption on rutile (1 1 0) in water adopting different a priori surface descriptions. In fact, the peptides were assumed to interact with a non-hydroxylated surface, corresponding to the surface observed in vacuo, hence exposing only 2-fold coordinated oxygens and 5-fold coordinated Ti atoms (Carravetta and Monti, 2006; Monti et al., 2008), whereas for the lipids adsorption both hydroxylated and non-hydroxylated surfaces were considered (Fortunelli and Monti, 2008). In this context we note again that the assumption of a non-hydroxylated surface is consistent with the experimental observation that water adsorption on this surface is mostly molecular and non-dissociative (Diebold, 2003), unlike that predicted by DFT methods. In view of the these remarks and of the simulations just mentioned, that will be commented upon in the next paragraph, we recently started a theoretical study of the surface behaviour of the TiO2 polymorphs using classical MD simulation techniques, adopting as a first step neutral, non-hydroxylated surfaces of the TiO2 polymorphs stable at ambient pressure, namely rutile, anatase and brookite (Raffaini G and Ganazzoli F, unpublished results). The main purpose of this preliminary study was to assess the possible effect of surface topology at the same surface chemistry in terms both of the exposed atomic species and of their coordination geometry. In our approach, as a proof of concept, we modelled (see Section 5.6: Appendix A) the interaction of bulk water with a TiO2 surface considering the three polymorphs exposing the simplest low-index surfaces with only 2-fold coordinated bridging oxygen atoms and 5-fold coordinated Ti atoms as in vacuo. These preliminary simulations on model surfaces with the surface atoms in the same local coordination, but in different arrangements due to the dissimilar crystal symmetry clearly indicate a different surface hydration as a purely

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5.3 A snapshot of the distribution of the water molecules on: (a) rutile (0 0 1), (b) anatase (1 0 0), and (c) brookite (1 0 0) viewed parallel to the surface at right angles.

topological effect. In fact, our results show a large ordering of the water molecules in a few ordered hydration layers on rutile (0 0 1), a smaller ordering on anatase (1 0 0) and an essentially disordered arrangement on brookite (1 0 0), as visually shown by Fig. 5.3. A more quantitative description of the statistical distribution of the water molecules is provided by the pair distribution function (or PDF for short) calculated from the MD trajectories. The PDF gives the probability density of finding a given set of atoms (for instance, in this case, the water oxygen atoms) as a function of the distance r from another set of atoms (the exposed surface atoms, here). The PDF plots obtained from the MD runs of the water molecules on the surfaces shown in Fig. 5.3 are reported in Fig. 5.4. The water molecules are held close to the surface by hydrogen bonds with the exposed bridging oxygen atoms, which gives rise to the first peak at a distance r Ê in the PDF curves of Fig. 5.4 for all the surfaces. Incidentally, we notice of 3.3 A that because of this interaction the hydrogen atoms of the first water layer are largely oriented towards the surface and closer to it than the oxygen atoms. Moreover, the greater height of the first peak indicates a larger ordering on the rutile surface compared to the other polymorphs, with relatively low mobility of the water molecules. We also point out that on the selected surfaces there is not enough room to have a significant amount of physisorbed water on the exposed Ti atoms, so that no peak is present at shorter distances, unlike what happens for the rutile (1 1 0) surface (Zhang et al., 2004). Further peaks in the plots in Fig. 5.4 at larger r

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5.4 The pair distribution function of the water oxygens as a function of their distance from the surface atoms on the TiO2 polymorphs (see text).

Ê values indicate additional hydration shells. In particular, the peak at about 6 A and those at larger distances indicate a second and possibly other hydration shells above to rutile surface, as also clearly seen in Fig. 5.3. On the anatase surface, the water molecules show a somewhat smaller local ordering compared Ê. to rutile, since only a second hydration shell is evident through the peak at 5.2 A Finally, the brookite surface only shows short-ranged water ordering in the first Ê , being essentially featureless at larger distances, as also hydration shell at 3.3 A shown by Fig. 5.3. Owing to the limited choice of the surfaces in these preliminary runs no definite conclusions can be drawn about the macroscopic wettability of the TiO2 polymorphs as determined, for instance, by contact angle measurements. However, these preliminary results strongly indicate a significantly different hydration pattern of the surface of the TiO2 polymorphs as a purely topological effect on the sub-nanometre scale, a pattern that in turn may affect the physico-chemical behaviour of the surfaces. We conclude this section by noting that the above-mentioned details of the surface hydration, also including the water orientation at the interface, or the adsorption of small molecules in general can be experimentally studied with the vibrational sum frequency spectroscopy, or VSFS, which can efficiently detect the vibrational features at the interface, basically unaffected by the contribution of bulk water (Richmond, 2002). Very few papers dealt with the water structure at a TiO2 surface, considering very thin films structurally not characterized (Kataoka et al., 2004; Uosaki et al., 2004). The results indicate strong sensitivity of the observed spectra as a function of pH (Kataoka et al., 2004), while the water hydrogens are oriented towards the surface, in keeping with our simulation results.

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Adsorption phenomena on TiO2 surfaces: classical MD simulation studies of the adsorption of larger molecules and of proteins

We now turn our attention to the MD simulations of mono- and dipeptides on the rutile (1 1 0) surface described at length in the previous paragraphs (Carravetta and Monti, 2006; Monti et al., 2008). These simulations were carried out in explicit water, but quantum DFT calculations in vacuo on a single amino acid (ionized alanine with a carboxylate anion) were also performed as a test case (Carravetta and Monti, 2006). The latter calculations showed a bidentate interaction of the carboxylate anion on two adjacent 5-fold coordinated Ti atoms as the preferred interaction geometry. The MD simulations carried out in water at a temperature of 310 K did confirm such interaction geometry, driven by the electrostatic interaction between the COOÿ anion and the formally Ti3+ atoms. In fact, a weaker interaction was obtained when the carboxylate anion was replaced by an amide group, for instance: in this case, the interaction takes place through the amide CO group and the amino NH2 group of alanine on adjacent Ti atoms. Minor differences in the geometrical parameters such as a larger value for the bidentate Ti


O distances in the MD simulations in water compared to the quantum DFT calculations in vacuo are likely due to hydration effects. Further simulations were also carried out for two dipeptides carrying zwitterionic ends, namely Ala-Glu, with a net charge of ÿ1 due to a COOÿ side group, and Ala-Lys, with a net charge of ‡1 due to a NH3+ side group. In both cases, the dipeptides showed a bidentate coordination geometry onto adjacent Ti atoms through the terminal COOÿ group of alanine. Such coordination was very strong and remained unchanged throughout the simulation, so that no major molecular rearrangement was observed (Carravetta and Monti, 2006). This study assumed a priori a well-defined adsorption geometry for the charged dipeptides as a starting point for the MD simulations, and therefore few overall changes were found due to the reasonable, but possibly biased, hypothesis of the initial bidentate carboxylate coordination to adjacent Ti atoms. In a subsequent paper of the same group (Monti et al., 2008) the bias was removed and the interaction of the same dipeptides in different starting orientations on the same rutile (1 1 0) surface was investigated in water, assuming, however, completely uncharged molecules. Among the arrangements that remained physisorbed, the most favourable interaction geometries obtained after the MD runs involved either multiple interactions of the CO amide group and NH2 groups with three neighbouring Ti atoms, or intermolecular hydrogen bonds with bridging oxygen atoms. Optimization of these surface interactions involved intramolecular rearrangements improving the molecular hydration and the surface interactions. Only in one case was the carboxylic COOH group found close to the surface, unlike what was found in the presence of the carboxylate anion, indicating the relevance of charged groups on the interaction geometry. In

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fact, a zwitterionic form of the Ala-Glu dipeptide was indeed again found to give a bidentate coordination of the COOÿ anion to two adjacent Ti atoms, consistent with the experimental results of X-ray photoelectron spectroscopy (Monti et al., 2008), while the terminal NH3‡ group formed intermolecular hydrogen bonds with the bridging oxygen atoms. In spite of the simplifications adopted of neutral, or at most zwitterionic dipeptides, these atomistic MD simulations suggested the possibility of surface interaction through the COOÿ group and, to a lesser extent, through the CONH group at the exposed 5-fold coordinated Ti atom, with possible hydrogen bonds involving neighbour bridging oxygen atoms. In view of these results, we have also begun to model the interaction of some subdomains of globular proteins on the same neutral, non hydroxylated surfaces of the TiO2 polymorphs rutile, anatase and brookite previously studied for their hydration using again atomistic simulations (Raffaini and Ganazzoli, unpublished results). We report here some preliminary results of this ongoing study, carried out with the same protein subdomains we considered in previous work to model protein adsorption on a hydrophobic graphite surface and on a hydrophilic poly(vinyl alcohol) surface, as recently reviewed (Ganazzoli and Raffaini, 2005; Raffaini and Ganazzoli, 2007). These protein subdomains have an unlike secondary structure: a human serum albumin subdomain formed by 60 amino acids comprising only -helices (Raffaini and Ganazzoli, 2003), and a fibronectin subdomain with two type I modules with 93 amino acids forming only antiparallel -sheets (Raffaini and Ganazzoli, 2004a). The simulations were carried out following the same simulation protocol adopted in previous work (Ganazzoli and Raffaini, 2005) and briefly summarized before for our results about the surface hydration. However, for computational reasons in this case we adopted an implicit solvent by using a distance-dependent dielectric constant instead of explicit water. Different starting geometries mimicking a random approach from solution were chosen for each subdomain on the TiO2 polymorph surfaces, namely rutile (0 0 1), anatase (1 0 0) and brookite (1 0 0). In this way, all the main sides of each subdomain can interact with the surface. The energy minimization of the system yielded the geometry in the initial adsorption stage shown in Fig. 5.5. In this stage, the molecules display a favourable interaction with the surface, leading to some minor spreading and flattening on the surface. Considering the various orientations of the subdomains approaching the surface from the bulk solution, we can try to correlate the interaction energy with the number of the residues that are in contact with the surface for each orientation, as done in Ê as the distance for a contact interaction on a TiO2 previous work. Taking 6 A surface, the resulting plot is shown in Fig. 5.6. The figure shows that the interaction energy, defined as the energy required to desorb the protein subdomains from the surface and bring them back to the free native state, increases linearly with the number of residues in contact with the surfaces, as previously found for

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5.5 The most stable initial arrangements of the two albumin subdomains (first row) and of the fibronectin models (second row) on the selected TiO2 surfaces: on rutile (left column), on anatase (central column) and on brookite (right column). In each case, a side view and a view from the above are shown for the adsorption on a finite surface. The secondary structure of the subdomains is shown through cylinders for the albumin -helices, and through very short arrows for the fibronectin -sheets.

5.6 The interaction energy Eint of the protein subdomains interacting with the TiO2 surfaces (see text) in the initial adsorption stage as a function of the number of residues that are in contact with the surface at a distance of 6 Ð, n6Ð. The filled and empty symbols respectively refer to albumin subdomain and the fibronectin modules in different orientations. The straight lines are the bestfit lines through the origin for the protein subdomains on rutile (solid line, r ˆ 0:9338), on anatase (dashed line, r ˆ 0:9115) and on brookite (dotted line, r ˆ 0:8649).

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other surfaces (Raffaini and Ganazzoli (2007) and citations therein). The slope of the best-fit straight lines through the origin gives the interaction energy per residue in contact with the surfaces: the values are 34:2  1:2 kJ/mol on rutile, 39:4  1:5 kJ/mol on anatase, and 32:2  1:6 kJ/mol on brookite, where the  sign indicates the estimated standard error of the fit. These values show that on average the adsorption is stronger on anatase (1 0 0) than on the chosen surfaces of the other polymorphs, even though the rutile-anatase difference is statistically significant only at the 2.8  level. There is a significant scatter in the calculated points, since when different sides of the protein subdomain interact with the surface there may be an unlike ratio among the number of hydrogen bonds and of the dipolar and dispersive interaction due to the exposed amino acids. For this reason, the adsorption of the albumin subdomain is slightly favoured compared to the adsorption of the fibronectin modules, on a per residues basis, even though the difference is generally minor and statistically not very significant. On the other hand, the larger size of the fibronectin module leads to an overall stronger interaction compared to the albumin subdomain. In any case, these findings show again the importance of the surface nanotopology, even at the same surface chemistry. The optimized arrangements in this initial adsorption stage show a small spreading and flattening on the surface, amounting to a few percent with respect to the principal axes of the native geometry, in terms of the components of the radius of gyration tensor parallel and perpendicular to the surface. The arrangements shown before in Fig. 5.5 can correspond to the final adsorption stage at a large surface coverage hindering the molecular spreading because of the smaller adsorption area available. On the other hand, a larger spreading could be expected for isolated molecules on a bare surface, a possibility that is currently being investigated by MD simulations at room temperature. On a bare surface, the molecules may show a stronger surface interaction due to a much larger number of residues interacting with the surface in particular with hydrophobic graphite (Ganazzoli and Raffaini, 2005; Raffaini and Ganazzoli, 2007). On the latter surface, in fact, the increase in the interaction energy was shown to be much larger than the energy cost imposed by the induced deformation due to the loss of the secondary structure and of the intramolecular hydrogen bonds. Preliminary results suggest that on the TiO2 polymorphs this is also the case, and additionally that the ordered structure of the crystalline surface might also drive some nanopatterning of the adsorbed subdomain possibly akin to what observed for instance on graphite (Raffaini and Ganazzoli, 2004b). The detailed results of these simulations will be published separately in due time. In the meantime, the adsorption of lipids on TiO2 in water was also investigated by MD simulations, considering hydroxylated, non-hydroxylated and partially hydroxylated rutile (1 1 0) surfaces (Fortunelli and Monti, 2008). The hydroxylated surface was assumed to show both TOH (terminal OHs on the 5-fold coordinated Ti atoms) and BOH hydroxyls (protonated bridging oxygen atoms). The modelled

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lipids were a (zwitterionic) phosphatidylcholine and a (anionic) phosphatidylserine derivative of the oleic acid (DOPC and DOPS, respectively), and a (cationic) myristic acid derivative of trimethylammonium propane (DMTAP), used as building blocks to form solid-supported phospholipid bilayers on SiO2 and TiO2. Molecules of each lipid were placed on the non-hydroxylated and the partially or fully hydroxylated surfaces, so that the charged head group was very close to the surface above the Ti atoms or the exposed OH groups. The results of the MD simulations indicated a strong coordination on the non-hydroxylated surface due to the interaction between the charged phosphatidylcholine and the exposed Ti atoms (formally a Ti3‡ cation) through a phosphate oxygen, assisted by a similar interaction of a neighbouring carbonyl group in DOPC, or through the carboxylate anion of the phosphatidylserine moiety of DOPS forming a bidentate interaction with two adjacent Ti atoms. On the other hand, the bulkier trimethylammonium propane head group of DMTAP led to a looser interaction with the bridging oxygen atoms. In the fully hydroxylated surfaces the charged head groups of DOPC and DOPS cannot directly interact with the Ti atoms shielded by TOH hydroxyls: therefore, the lipid molecules displayed a significant mobility in the MD runs with large fluctuations leading also to temporary detachment of DOPC from the surface and to an almost complete detachment of DMTAP. The partially hydroxylated surface showed a somewhat intermediate behaviour, but since some Ti atoms are exposed to the environment, the lipids can become strongly adsorbed on these sites, as indeed observed experimentally in particular for DOPS. In addition to a study of the local details of the head group interaction with the surface atoms, the MD simulations also provided information about the fluctuations and hinge-bending motion of the anchored molecules and their higher-temperature surface diffusivity, showing the array of information that can be potentially achieved by atomistic MD simulations.

5.3

Technologies for tailoring Ti oxides on titanium

As previously discussed, the thin titanium dioxide layer differently structured and oriented, always covering the surface of titanium and its alloys, has the potential to drive different interface reactions with the many possible environmental molecules. From a practical point of view, the nature and quality of the oxide film covering titanium medical devices is strongly affected by the mechanical, thermal and sterilization conditions they have been submitted to during the manufacturing process. A TiO2 surface is chemically stable (VoÈroÈs et al., 2001): no significant modifications on its thickness, for instance, have been noticed after one year of exposition to ambient atmosphere (Sittig et al., 1999a). However, a titanium dioxide surface is not inert when in contact with a physiological environment, and the chemico-physical processes explained in the previous section drive the surface interactions with the water and the biological molecules. For this main

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reason, together with the fact that the currently used ISO standard titanium grades and alloys meet the requirements for safe, long-term performance in the majority of medical devices (VoÈroÈs et al., 2001), the fundamental `playground' of biomaterial scientists working on titanium surfaces for medical applications is represented by surface modulation of chemical composition and structure of the native oxide layer, in order to promote specific biological response. Surface modification technologies exploited in biomedical applications have been accurately reviewed by some authors (Liu et al., 2004) and can be schematically classified, regarding TiO2 modification, as: 1. Mechanical modification treatments, in general aimed at obtaining specific surface topographies via subtraction processes (i.e. blasting, grinding, polishing) in the range of m to mm scale (removed from the material); these methods generally do not allow any modulation of chemical and structural composition of the obtained oxide layer, even if they can result in residual stresses and plastically deformed oxide layer (Lausma, 2001). 2. Chemical modification treatments, aimed at: (a) Removing the oxide layer, descaling or etching the surface, via subtraction processes (i.e. acid treatments, electrochemical etching and polishing). Rough or smooth surfaces and clean surfaces are achieved in the range of m to mm scale (removed from the material) and the resulting oxide surface is in general a native-like oxide surface in terms of chemical structure and composition, depending on the exposed environment; (b) Coating deposition of inorganic or organic films, with alteration of the surface topography (i.e. electrochemical treatments, CVD, biochemical methods, sol gel). Surface coatings with suitable topographic characteristic are achieved by deposition of films with a tunable morphology, chemical composition and structure. 3. Physical modification treatments, aimed at: (a) Removing oxide or etching the surface, via subtraction processes (i.e. sputtering, glow discharge plasma treatments). Rough or smooth and clean surfaces can be achieved, in the range of nano- to several micrometre scale. Moreover, they are used to clean, sterilize and remove the native oxide layer; (b) Surface functionalization (i.e. ion implantation techniques). A modification of surface composition in the range of nano- to micro-metre scale; (c) Coating deposition (i.e. thermal spray, PVD, ion implantation, glow discharge plasma treatments). These technologies can also be coupled and associated to exploit the specific effects held by each one, with the aim of obtaining a complex surface and structure on a medical device. However, in an overall perspective, chemical and, in particular, electrochemical technologies are, in our opinion, of particular interest due to the following reasons:

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· they find wide application in industrial production of medical devices ± allowing a simple scale-up in case of significant advancement in basic research; · they provide efficient and uniform access to all surfaces, even on complex 3D devices such as screws and cardiovascular stents (Variola et al., 2009); · they open the possibility for advanced surface modifications, with a specific modulation of the oxide layer onto materials and geometries commonly used in biomedical devices.

5.3.1

Chemical modifications of structure and composition of native TiO2

Chemical surface modification methods are based on chemical reactions that can occur at the interface between titanium and a suitable reagent. It is one of the most exploited technologies, since several chemical surface treatments are currently used in commercial biomedical applications, a large part of these treatments often being proprietary (Lausma, 2001). This class of treatments ranges from simple decontamination procedures to more complex coatings. Cleaning procedures Solvent-based surface decontaminations are in general not intended for TiO2 structure modification, even if correct cleaning is mandatory before sterilization and clinical use and represents a key point when studying interface processes of native oxide layers, having a strong influence on in vitro and in vivo responses. In fact, the TiO2 layer shows a significant reactivity with organic material, such as insertion reactions or complex formation (Textor et al., 2001). When a reaction of an alcohol with a R ˆ hydrocarbon chain and titanium oxide surface takes place, two main species are formed at the material surface: · Alkoxides Ti±OH + HO±CH2±R ÿ! Ti±O±CH2±R + H2O

5.1

· Peroxides Ti±OOH + HO±CH2±R ÿ! Ti±OO±CH2±R + H2O

5.2

Analogously, a reaction with caboxylic acid leads to the formation of carboxylates: Ti±OH + HOOC±R ÿ! Ti±OOC±R + H2O

5.3

Chemisorption and physisorption of hydrocarbons also take place. This may strongly affect the surface wettability (see, for instance, Wang et al. (1997) and Thompson and Yates (2006)). Even if several decontamination processes have been described, their selection strongly depends on the substrate and on the nature of contaminations to be

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removed and no exclusive indication can be provided. However, no matter how carefully a titanium surface is cleaned, it will always show a significant carbon signal on spectroscopic analyses (Lausma, 2001). These contaminants strongly affect the overall response of the in vivo implants, both on cell differentiation and growth factor production (Schwarz et al., 2009), and should be taken into account when evaluating the chemico-physical response of the oxide to biomolecules. Wet chemical etching Wet chemical etching is a standard procedure for biomedical surface preparation, commonly performed on implantable prostheses (e.g. dental implants) and aimed at surface decontamination and morphology modification. Technologies used to manufacture metallic implants and devices, such as forging, machining, casting always induce some contamination of the surface. At an industrial level the finishing of an implantable metal device must be processed to detach and to eliminate any contaminant sitting or embedded in the first surface layer. The chemical stability of titanium dioxide limits the number of solutions that can be used to treat these surfaces; nevertheless, this is the most popular and efficient way to modify surface at the nanoscale and it is attractive for large-scale manufacturing (Variola et al., 2009). Even if the number of effective chemicals is limited by the reactivity of titanium dioxide, acid and alkaline water solution-based approaches are possible (Lausma, 2001; Wennerberg and Albrektsson, 2009): this class of treatments leads to the reaction with the oxide layer, mainly resulting in specific morphological and chemical modifications. Acid etching operated in proper solutions can be effectively and economically used to dissolve the protective oxide layer and to etch the metal substrate. The acid-based approach is based on the use of HF (MaÂrcia Soares et al., 2005), HCl, H2SO4, malonic acid, and selected mixtures of these acids with HNO3 and H2O2. Many etching surface treatments were developed to improve osseointegration of implantable titanium devices, especially for dental applications, and today wet etching of titanium is widely diffused. Titanium and its alloys can be effectively etched in fluoride containing solutions (capable of etching and dissolving the highly stable and protective titanium oxide layer), with different acid mixtures. The oxide and metal dissolution can also be operated to provide an increased roughness to the surface, contributing to improving hard tissue fixation through mechanical interlocking of the device surface to the bone. After the etching process, a new protective oxide film can easily be reconstituted in air or in another oxidizing environment. An example of such a titanium surface, obtained in our laboratories through a two-step chemical etching process, is shown in Fig. 5.7.

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5.7 Chemical etched titanium surface through a two-step process: (1) a HF etching process is performed to activate the titanium surface (passive TiO2 film removal); (2) high temperature HCl and H2SO4 etching for a few minutes provides the microrough surface.

Wet chemical oxidation Oxidation of TiO2 surfaces has been exploited mainly by using two strong oxidants, H2O2 and HNO3. HNO3 solutions (20±40% by volume) are common standard passivation procedures for titanium and titanium alloys and often represent the last step in a surface preparation process, but have no major influence on the overall surface topography of a titanium surface. Passivation with nitric acid does not change the overall thickness of the oxide film, but it could improve the homogeneity of the passive film at local defects (Sittig et al., 1999b). H2O2 surface treatment approach is particularly interesting because some studies on protein adsorption, that can be correlated to theoretical calculations, have been performed (Sousa et al., 2007). Sousa et al. have shown the kinetic behaviour of Fibronectin adsorption by its radiolabelling, XPS, and ellipsometry on TiO2 formed on commercially pure titanium after immersion in H2O2 (TiO2 cp in the text) and TiO2 sputtered on Si (TiO2 sp in the text). Chemical treated surface adsorbed more Fibronectin than sputtered surface, after 60 min of adsorption. Fibronectin molecules do also adsorb more strongly to the former surface. The aggregate structure had an intermediate feature shared by some protein fibrillar assemblies at interfaces, which is believed to promote cell adhesion and cytoskeleton organization (Sousa et al., 2007). Similarly, in a

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previous study, the same group (Sousa et al., 2004) studied the kinetics of adsorption of albumin on the same surfaces, showing that TiO2 cp adsorbs less albumin than TiO2 sp, as shown by the radiolabelling data, but the adherent molecules are strongly attached to the former surface, as indicated by the work of adhesion and by the desorption studies. Moreover, they showed that at larger concentration albumin adsorbs as a multilayer. Furthermore, competitive adsorption measurements (Sousa et al., 2008) indicate that a significantly larger amount of fibronectin is adsorbed on TiO2 cp or on TiO2 sp than human serum albumin, at least when the two proteins are present in a 1:1 ratio. On the other hand, larger amounts of adsorbed albumin are observed at the physiological ratio due to the much larger abundance, even though this amount is not as large as from a single protein solution, being effectively limited by fibronectin. These findings suggest a stronger binding to the surface by fibronectin, in qualitative agreement with the calculated behaviour previously reported. Sol-gel coatings Sol-gel is a synthetic technology for the synthesis of inorganic materials from chemical routes and is a powerful tool for TiO2 thin coating deposition and structuring. This technology finds a widespread application: several interesting industrial products have been produced, such as `IROX', an antireflective coating commercialized since 1980s by Schott. Even if sol-gel is a process commonly used for silicate preparation, systems based on transition metals (e.g. V, Zr, Ti) can be effectively produced: the synthetic steps and the involved solution chemistry are quite similar, since a greater chemical reactivity is associated with transition metals, which exhibit several coordination states together with a lower electronegativity when compared to Si alkoxides (Brinker and Scherer, 1990). Sol-gel involves a sequence of chemical and physical steps that allow bulk ceramics to be obtained, starting from chemical precursors; this sequence is schematically based on: (i) inorganic polymerization (based on hydrolysis and polycondensation), (ii) gelation, (iii) aging, (iv) drying, and (v) densification and crystallization. This complex theoretical approach is actually quite simple, provided that the chemical synthesis and control of process parameters are well set up. For transition metals, however, the lower stability toward hydrolysis, condensation and nucleophilic reactions compared to Si alkoxides requires a stricter control of processing moisture and hydrolysis conditions. In biomedical applications, sol-gel processes have been exploited to produce silica-derived bioactive materials in thin film and monolith preparation, especially those synthesised in binary (SiO2±CaO) and ternary (SiO2±CaO± P2O5) systems (Arcos et al., 2009). As far as titanium oxide is concerned, thin coating preparation on different substrates, including titanium and titanium alloys, is possible.

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From a chemical point of view, the first step (inorganic polymerization) can be based on two synthetic routes, either starting from metallo-organic precursors (such as alkoxides Ti(OCnH2n+1)4) in nonaqueous solvents or inorganic salts (e.g. TiCl4) precursors in aqueous solution, even if the former is the more commonly used and described process. Alkoxides are in general dissolved in their parent alcohol, and hydrolysis is promoted by using water, according to the following hydrolysis-condensation reactions: · Hydrolysis: Ti±(OR)4 + H2O ÿ! HO±Ti±(OR)3 + R±OH

5.4

· Condensation: (OR)3Ti±OH + HO±Ti±(OR)3 ÿ! (OR)3±Ti±O±Ti±(OR)3 + H2O (Oxolation)

5.5

(OR)3Ti±OR + HO±Ti±(OR)3 ÿ! (OR)3±Ti±O±Ti±(OR)3 + R±OH (Alcoxolation)

5.6

These competitive reactions lead to the formation, in a controlled environment of a colloidal suspension of solid nanometric particles (sol). These particles tend to aggregate either spontaneously or by introducing chemical and physical instabilities, eventually resulting in a continuous 3D skeleton network, interpenetrated by a continuous liquid phase (gel). When thin films have to be deposited on a substrate, the gelation process is generally achieved during the coating deposition stage, that can be realized via conventional film formation techniques, such as spin-, dip-, or spray-coating. Electrophoresis is also possible, in a way that is described later in this chapter. After the gel formation, the liquid is removed from the interconnected pore network. The final step is generally constituted by heating the porous gel that results in a densification of pores, leading to a compact and dense film: the densification temperature depends considerably on the dimensions of the pore network, the connectivity of the pores, and the surface area. This last step is challenging in tuning the crystallographic form of the titanium oxide: the calcination process may, in fact, be applied if a denser and more crystalline oxide layer is desired (Advincula et al., 2006). TiO2 sol-gel coatings, based on either conventional (Gupta and Kumar, 2008) or Stepwise Surface process (Advincula et al., 2006) have been the subject of several studies aimed at understanding their effects on water interaction, calcium phosphate (Ca/P) crystallization and protein absorption (Advincula et al., 2006; DieudonneÁ, 2002; Li and de Groot, 1993; and citations therein). Sol-gel-derived TiO2 coatings have been shown to chemically bond to bone after implantation (Li and de Groot, 1993): this early evidence has spurred following studies aimed at understanding the bioreactivity of such surfaces on titanium alloys. For instance, Advincula et al. (2006) showed that TiO2 sol-gel

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coatings, if compared to passivated surfaces, demonstrate an increased capacity to initiate Ca/P nucleation. This behaviour has been explained by the presence of Ti-OH species on sol-gel surfaces, which promote electrostatic attraction of Ca++ or hydrogen bonding with phosphates. Moreover, the presence of hydroxyl groups and a specific crystallographic surface influences the deposition of proteins on TiO2 sol-gel coatings and their presence (of both albumin and fibrinogen) in a simulated body fluid (SBF) solution delays the growth of calcium phosphates on the surface of TiO2 coating through decreasing the recrystallization rate of the initially formed amorphous calcium phosphate (Areva et al., 2002). According to this study, a thin layer of amorphous Ca/P layer adsorbs in the early stage of immersion in SBF; when proteins (either albumin or fibronectin) are present in solution, they tend to subsequently adsorb. Both water molecules and ions in solution play an important role in modulating the protein adhesion and surface conformation; moreover this work suggests that adsorbed proteins have an influence in Ca/P nucleation and growth kinetics. Theoretical modelling of these complicated systems involving the competitive adsorption of proteins from a solution in the presence of different electrolytes is a challenging task that has begun to be addressed to some extent only very recently. We can confidently expect, however, that in the near future computer simulations will offer new insights about these phenomena, complementing the experimental observations and providing a deeper insight for the design of biomaterial surfaces.

5.3.2

Electrochemical modifications of structure and composition of native TiO2

Electrochemical surface modifications on titanium and its alloys can be performed in three different ways (Fig. 5.8): · Anodic oxidation, in which the surface to be treated is exposed to an electrolytic bath as the anode and TiO2 film grown as a consequence of metal oxidation and oxygen diffusion processes (Fig. 5.8a). · Cathodic polarization, sometimes referred as electrolytic deposition (ELD), in which the surface to be treated is exposed to an electrolytic bath as cathode, and inorganic matrix is deposited over the surface as a consequence of a local pH variation, ion transport and species reduction (Fig. 5.8a and b). This process is generally exploited in calcium phosphate deposition. · Electrophoretic deposition (EPD), in which the surface to be treated is immersed in an organic solvent-based electrochemical cell as cathode or anode, and the coating is grown as a consequence of the deposition of suspended charged powder particles onto a conductive substrate of opposite charge, by applying a DC electric field (Fig. 5.8b).

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5.8 Schematic of (a) electrochemical anodic and cathodic surface modifications (Yerokhin et al., 1999) and (b) cathodic electrophoretic deposition (EPD) vs. electrolytic deposition (ELD), showing electrophoretic motion of positively charged ceramic particles and ions followed by hydrolysis of the ions to form colloidal nanoparticles and coagulation of the particles to form EPD and ELD deposits (Zhitomirsky, 2002).

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Anodic polarization and anodic spark deposition Anodic polarization is an electrochemical technology that allows thick and opportunely structured metal oxide films over the surface of a metal to be obtained. The device is connected to the anode of a power supply and immersed in an electrolytic bath: electrode reactions in combination with electrical fielddriven metal and oxygen ion diffusion lead to the formation of an oxide film at the working surface. The explanation of the electrochemical processes involved in anodization can be found in several works of the early 1980s: based on the use of implanted radioactive noble gas atoms as markers (Hurlen and Gulbrandsen, 1994), it has been shown that both metal and oxygen migration generally contribute to the charge transport occurring in growing anodic films. The relative importance of these transport processes depends on the metal and the current density more than on the anodizing solution, provided it does not give rise to pore formation (Hurlen and Gulbrandsen, 1994). From a physicochemical point of view, anodic oxidation reactions of titanium surfaces depend on interface processes. At the Ti/Ti oxide interface, the main reaction is represented by the oxidation process: Ti „ Ti2‡ + 2eÿ

5.7

At Ti oxide/electrolyte interface: 6H2O „ 2O2± ‡ 4H3O+

5.8

6H2O „ O2 " ‡ 4H3O‡ + 4eÿ

5.9

At both interfaces: Ti2+ + 2O2± „ TiO2 + 2eÿ

5.10

The growth of an oxide film is then a consequence of oxygen diffusion together with interface redox processes. Anodization is hence used to produce films of increased oxide thickness, porous coatings and selected crystallographic forms. More recently, TiO2 nanotubes have also been proposed by several authors (Oh et al., 2005). A scheme of the voltage/time evolution of the process at steady current is reported in Fig. 5.9a: during anodization, that can be conducted either in galvanostatic or potentiostatic mode, anodic oxide film growth carries on as long as the electric field is strong enough to drive appropriate chemical species through the film. Two different stages of the anodic oxide film growth can be evidenced. During the first phase, TiO2 film increases its thickness and the anodic growth is almost linearly dependent on the applied voltage. Voltage drop in the electrochemical cell is mainly correlated to the oxide film thickness, because it exhibits higher resistivity than other electrochemical cell elements (e.g. bath composition, counter electrode). Growth of such a dielectric film is linear up to

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5.9 (a) Voltage-time plot of an anodic oxidation process performed on c.p. grade 2 Ti at steady state current density 70 A mÿ2: processes at high voltages result in formation of a micro-porous TiO2 film, with the insertion of ionic species composing the electrolytic bath (courtesy of the authors from Chiesa et al., 2003) and (b) Oscillographic observations of the sparks occurring during the galvanostatic anodization of titanium at high current density (Delplancke and Winand, 1988a).

100±160 V, depending on process conditions (Chiesa et al., 2003). When anodization is carried out at higher voltages (second stage), the process leads to an increased gas evolution and often to sparking. In these conditions, micro-arc oxidation (MAO) or anodic spark deposition occurs (ASD): MAO is an electrochemical technique inducing the formation of thick and porous ceramic coatings

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on valve metals (titanium, tantalum, etc.). By using these technologies, physicochemical, morphological and structural features of oxides on titanium can be tuned over a wide range, namely thickness, ion insertion and crystallographic TiO2 structuring. Ion insertion via low voltage anodization is a controversial issue: according to some studies, anodic film composition is constituted mainly by titanium dioxide (Lausmaa et al., 1990). However, other works, based on XPS analysis or Rutherford backscattering reported oxygen enrichment, embedding of electrolyte elements in the oxide film and the presence in the very outer oxide layer of a certain amount of hydrogen (Aladjem, 1973; Ask et al., 1989; Lausmaa et al., 1990; Delplancke and Winand, 1988a,b). Ti6Al4V behaves very similarly to pure graded titanium, although higher concentrations of Al relative to titanium in the oxide than in the metal and depletion of vanadium in the outermost layers have been observed (Liu et al., 2004). Different kinds of structural defects, such as pits and sometimes pores, were often detected (Delplancke and Winand, 1988a,b; Larsson et al., 1994). Diffraction studies showed the oxide films to be either totally amorphous or in some cases partially crystalline. The most commonly observed phases are anatase and rutile (both tetragonal), but also Ti3O5 and brookite (orthorhombic or tetragonal) have been detected (Aladjem, 1973). This aspect is of particular interest because, in recent works of our research group, the presence of anatase phase has been demonstrated to play an important role in surfaces diminishing the bacterial colonization (Del Curto et al., 2005). As previously discussed, when anodization is carried on at higher voltages micro-arc oxidation occurs, as a consequence of an increased gas evolution and generation of sparks spreading all over the surface. In Fig. 5.9b the behaviour of the oscillographic observations of the sparks occurring during a galvanostatic anodization of titanium is presented, as studied by Delplancke and Winand (1988a). Simultaneous sparking and ionic diffusion processes result in three main effects on MAO-treated surfaces. First of all, the sparking process promotes local melting and re-crystallization of the oxide film, generally resulting in the incorporation of ionic species with a concentration gradient along the oxide. The electrolytic bath composition affects the final chemical composition of the grown oxide (Yerokhin et al., 1999). Sparks generate a typical morphology with porosity at the micrometric scale, that can be usefully used in bone contact devices (Chiesa et al., 2003) and TiO2 significantly grown in thickness, determining an amelioration of the corrosion and ion release resistance properties (Yerokhin et al., 1999). Such morphological and chemical properties are achieved as a consequence of several variables, each one contributing to determine the coating properties and morphology: namely, the electrolyte bath composition, voltage and current density, temperature and process time. All these aspects have been reviewed

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with the aim of explaining the general aspects of plasma electrolysis (Yerokhin et al., 1999) and with a special emphasis on biomaterial applications (Chiesa et al., 2003). This process can be usefully applied for titanium modification and coating preparation for biomedical applications, exploiting both the morphological and chemical results of such a treatment. By MAO it is possible, for instance, to incorporate Ca, P and many other ions in the anodically grown oxide film, in order to promote the osseointegration process. Silver ions could be also considered to provide antibacterial properties to the metal surface. Moreover, these films show a controlled porosity, and are highly adherent to the substrate (Yerokhin et al., 1999). MAO-based treatments have been recently studied and considered to modify the titanium oxide morphology and composition, especially to improve osseointegration and fixation of dental implants. The biological response of MAO-based treatments showed very encouraging results (Chiesa et al., 2007), and some clinical applications are currently running. Figure 5.10 shows the morphology and EDS spectrum of a MAO-modified titanium surface developed at Politecnico di Milano. The typical micro-porous morphology can be clearly observed. Electrophoretic deposition (EPD) Electrophoretic deposition is an interesting electrochemical method for processing bulk and coating materials, mainly ceramics. EPD was discovered in the early 1800s and applied until the 1990s as a technology for conventional ceramic preparation; however, in the last 15 years, the interest in EPD as useful technology to produce advanced materials increased widely, both in academia and industry (Corni et al., 2008). Examples of electrophoretic deposition of almost any material class can be found, including metals, polymers, carbides, oxides, nitrides, and glasses: EPD appears of particular interest in surface modifications of materials for biomedical application, because it allows the deposit of ceramic and organic/inorganic coatings with high purity on complex geometries and shapes (Gottlander et al., 1997; Zhitomirsky, 2000; Zhitomirsky and GalOr, 1997). From a fundamental point of view, the mechanism of electrophoretic deposition involves two different steps (Moskalewicz et al., 2007): 1. Electrophoresis in which charged particles, in a suitable solvent, move toward an electrode of opposite charge when an electric field is applied between two electrodes. The deposition process can be either anodic or cathodic: by suitable modification of the surface charge on the particles, any of the two modes of deposition is in fact possible (Besra and Liu, 2007). Electrophoretic deposition can be hence applied to any solid available in the form of a fine powder or a colloidal suspension (Van der Biest and Vandeperre, 1999).

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5.10 (a) SEM and (b) EDS analysis of a Si, Ca, P, Na enriched titanium oxide developed by MAO to improve osseointegration of dental and orthopaedic implants.

2. Deposition on the working electrode, as a consequence of the fact that the motion of charged particles results in an accumulation of ceramic particles on it. Deposit formation is achieved via particle coagulation: a thorough understanding of the mechanisms of particle coagulation, crucial for successful electrodeposition of ceramic and organoceramic materials, has been proposed based on the classical Derjaguin±Landau±Verwey±Overbeek (DLVO) theory of colloidal stability (Zhitomirsky, 2002).

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The basic difference between EPD and electrolytic deposition/polarization is that the former is based on the suspension of particles in a solvent, whereas the latter is based on solution of ionic species, which in general undergo redox reactions leading to coating formation (Fig. 5.8). EPD process results in a compact powder coating, and therefore it should be followed by a densification step such as sintering or curing, in order to obtain a fully dense material (Van der Biest and Vandeperre, 1999), limiting in some cases its application. Such kinds of deposition have not been totally exploited in biomedical applications, but promising results, in terms of bioactivity and uniformity of obtained coatings, have been obtained (Gottlander et al., 1997; Zhitomirsky, 2000; Zhitomirsky and GalOr, 1997): however, an increasing interest can be noticed around this technology, as a useful tool for the surface treatment of titanium alloys. Moskaleicz et al. (2007), for instance, proposed the use of EPD as a tool for TiO2 crystallographic surface tuning: in this work, they were able to deposit a uniform and compact coating of TiO2 on Ti6Al7Nb, a two phase ( ‡ ) titanium alloy, used in orthopaedic prostheses as an alternative of grade 5 Ti alloy (Navarro et al., 2008), because it is free from vanadium. Ti alloy surface have been placed as a cathode in a solution of commercial TiO2 nanoparticles (Degussa P25, with a crystallographic composition of 80% anatase and 20% rutile) and a thick oxide film has been deposited. After annealing at high temperature, it was found via X-ray diffraction intensities corresponding to the anatase (1 0 1) and rutile peaks (1 1 0) ± i.e. the most thermodynamically stable phases of the two polymorphs, as previously pointed out ± that the fraction of rutile phase of oxide coating was approximately 10%. These results are particularly interesting, even though no studies on biological interfaces have been carried out: the possibility of simple tuning of the crystallographic structure of a TiO2 coating could be a useful tool to produce complex surfaces, in terms of treated geometries and relative ratios of crystallographic species.

5.4

Future trends

New applications for materials in life sciences are providing real challenges in terms of scope and complexity, perhaps not yet properly addressed by existing systems. Two main aspects constitute, in fact, the driving idea of modern biomaterials science: 1. Identification of specific design parameters `critical to performance': materials used in different fields can be transferred by analogy in medicine, then modified and optimized with the knowledge emerging from the studies of biological systems (Langer and Tirrell, 2004). 2. Design and tailoring of novel materials that exhibit `bio-like' behaviour enhanced with additional properties derived from abiotic synthesis or design (Alexander and Shakesheff, 2006).

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Advances in cell biology, chemistry, and materials science are enabling the production of a new generation of materials with multiple and combined functions inspired by biology. Such new biomimetic materials, including those with improved biocompatibility, specificity and other critical properties, have specifically been developed by applying modern methods of material science design and biomaterial science (Lendlein and Kelch, 2002). Likewise theoretical and simulation methods are providing the basic science interpretation to the surface reactions and interface phenomena. Atomistic computer simulations and/ or theoretical modelling are increasingly providing powerful access ± key to a fuller understanding. This approach has already been applied both at the quantum level to study surface reconstruction and reactivity, the details of adsorption coordination in chemisorption or in physisorption, etc., and at the classical level typically by molecular dynamics simulations to understand the surface adsorption at a chosen temperature of large molecules in systems comprising tens of thousands of atoms overall, to carry out a statistical analysis of the intramolecular surface rearrangements, the possible surface dynamics, and so on. Such theoretical approaches have already been applied to model the interface phenomena at TiO2 surfaces and we can easily predict that they will be increasingly pursued in the near future. In this context, we can expect that increasingly realistic systems will be modelled, considering not only the explicit solvent in the presence of appropriate electrolytes, but also increasingly sophisticated pictures of the TiO2 surfaces in terms of the specific crystallographic planes with an appropriate distribution of hydroxyls producing the effective surface charge densities experimentally observed at the pH of interest. Furthermore, the simultaneous presence of more crystalline phases (in particular rutile and anatase, as previously pointed out) poses further theoretical problems about the phenomena that may take place at the grain boundary interface and at their common interface with the outer environment. In this context, recent MD simulations (Deskins et al., 2007) have begun to model the rutile-anatase interfaces formed by the common surface of either polymorph, inter alia those among rutile (1 1 0) and anatase (1 0 1) (Moskaleicz et al., 2007). The simulations have shown the presence of a disordered interface having the thickness of about 0.4 nm, hence limited to the surface atomic layers, but most interestingly they may provide a basis to understand the influence of the interface on the mechanical and electronic properties of these polycrystalline materials, and on the surface adsorption phenomena at the grain boundary. At the same time, several surface modification technologies are emerging as powerful tools for tailoring TiO2. Among them, chemical and electrochemical techniques found widespread application in TiO2 chemistry and crystallography modifications on titanium alloys, aimed at promoting specific responses when in contact with biological environments. Several of these treatments are currently used in biomedical device industries, and an increasing demand, sometimes commercially driven, pushes forward their application. However, a better

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understanding of material/environment interaction fundamentals could help to tailor in a more specific and appropriate way these technologies and the obtained surfaces. In conclusion we confidently expect that an increasingly greater use of the theoretical and simulation methodologies may integrate and complement the experimental approaches providing a theoretical foundation of the observed phenomena about the performances of native and modified surfaces, eventually providing powerful help in the design of suitably tailored biomaterials.

5.5

References

Advincula M C, Rahemtulla F G, Advincula R C, Ada E T, Lemons J E and Bellis S L (2006), `Osteoblast adhesion and matrix mineralization on sol-gel-derived titanium oxide', Biomaterials, 27, 2201±2212. Aladjem A (1973), `Anodic oxidation of titanium and its alloys', J Mater Sci, 8, 688±704. Alexander C and Shakesheff K M (2006), `Responsive polymers at the biology/materials science interface', Adv Mater, 18, 3321±3328. Arcos D, Izquierdo-Barba I and Vallet-Regi M (2009), `Promising trends of bioceramics in the biomaterials field', J Mater Sci-Mater M, 20, 447±455. Areva S, Peltola T, SaÈilynoja E, Laajalehto K, LindeÂn M, and Rosenholm J B (2002), `Effect of albumin and fibrinogen on calcium phosphate formation on sol-gelderived titania coatings in vitro', Chem Mater, 14, 1614±1621. Ask M, Lausmaa J and Kasemo B (1989), `Preparation and surface spectroscopic characterization of oxide films on Ti6Al4V', Appl Surf Sci, 35, 283±301. Balamurugan A, Rajeswari S, Balossier G, Rebelo A H S and Ferreira J M F (2008), `Corrosion aspects of metallic implants ± an overview', Mater Corros, 59, 855±869. Barnard A S, Zapol P and Curtiss L A (2005), `Modeling the morphology and phase stability of TiO2 nanocrystals in water', J Chem Theory Comput, 1, 107±116. Besra L and Liu M (2007), `A review on fundamentals and applications of electrophoretic deposition (EPD)', Prog Mater Sci, 52, 1±61. Blunt L, Bills P, Jiang X, Hardaker C and Chakrabarty G (2009), `The role of tribology and metrology in the latest development of bio-materials', Wear, 266, 424±431. Brinker C J and Scherer G W (1990), Sol-gel science: the physics and chemistry of sol-gel processing, San Diego, CA, Academic Press. Brunski J B (1996), `Metals', in Ratner B D, Schoen F J and Lemons J E, Biomaterials Science: An Introduction to Materials in Medicine, Elsevier Academic Press: 37± 50. Cabailh G, Torrelles X, Lindsay R, Bikondoa O, Joumard I, Zegenhagen J and Thornton G (2007), `Geometric structure of TiO2(110)(1  1): Achieving experimental consensus', Phys Rev B, 75, 241403. Carp O, Huisman C L and Reller A (2004), `Photoinduced reactivity of titanium dioxide', Prog Solid State Chem, 32, 33±177. Carravetta V and Monti S (2006), `Peptide-TiO2 surface interaction in solution by ab initio and molecular dynamics simulations', J Phys Chem B, 110, 6160±6169. Charlton G, Howes P B, Nicklin C L, Steadman P, Taylor J S G, Muryn C A, Harte S P, Mercer J, McGrath R, Norman D, Turner T S and Thornton G (1997), `Relaxation of TiO2(110)-(1x1) using surface X-ray diffraction', Phys Rev Lett, 78, 495±498. Chiesa R, Sandrini E, Rondelli G, Santin M and Cigada A (2003), `Osteointegration of

ß Woodhead Publishing Limited, 2011

138

Surface modification of biomaterials

titanium and its alloys by anodic spark deposition and other electrochemical techniques: a review', J Appl Biomater Biomech, 1, 91±107. Chiesa R, Giavaresi G, Fini M, Sandrini E, Giordano C, Bianchi A and Giardino R (2007), `In vitro and in vivo performance of a novel surface treatment to enhance osseointegration of endosseous implants', Oral Surg Oral Med Oral Pathol Oral Radiol Endod, 103, 745±756. Cigada A, Cabrini M and Pedeferri P (1992), `Increasing of the corrosion-resistance of the Ti6Al4V alloy by high thickness anodic-oxidation', J Mater Sci-Mater M, 3, 408±412. Corni I, Ryan M P and Boccaccini A R (2008), `Electrophoretic deposition: From traditional ceramics to nanotechnology', J Eur Ceram Soc, 28, 1353±1367. Del Curto B, Brunella M F, Giordano C, Pedeferri M P, Valtulina V, Visai L and Cigada A (2005). `Decreased bacterial adhesion to surface-treated titanium', Int J Artif Organs, 28, 718±730. Delplancke J L and Winand R (1988a), `Galvanostatic anodization of titanium ± I. Structures and compositions of the anodic films', Electrochim Acta, 33, 1539±1549. Delplancke J L and Winand R (1988b), `Galvanostatic anodization of titanium ± II. Reactions efficiencies and electrochemical behaviour model', Electrochim Acta, 33, 1551±1559. Deskins N A, Kerisit S, Rosso K M and Dupuis M (2007), `Molecular dynamics characterization of rutile-anatase interfaces', J Phys Chem C, 111, 9290±9298. Diebold U (2003), `The surface science of titanium dioxide', Surf Sci Rep, 48, 53±229. DieudonneÁ S C, van den Dolder, de Ruijter J E, Paldan H, Peltola T, van `t Hof M A, Happonen R P and Jansen J A (2002), `Osteoblast differentiation of bone marrow stromal cells cultured on silica gel and sol-gel-derived titania', Biomaterials, 23, 3041±3051. Fortunelli A and Monti S (2008), `Simulations of lipid adsorption on TiO2 surfaces in solution', Langmuir, 24, 10145±10154. Freese H L, Volas M G and Wood J R (2001), `Metallurgy and technological properties of titanium and titanium alloys', in Brunette D M, Tengvall P, Textor M and Thomsen P, Titanium in medicine: material science, surface science, engineering, biological responses, and medical applications, Berlin; New York, Springer: 25±51. Ganazzoli F and Raffaini G (2005), `Computer simulation of polypeptide adsorption on model biomaterials', Phys Chem Chem Phys, 7, 3651±3663. Geetha M, Singh A K, Asokamani R and Gogia A K (2009), `Ti based biomaterials, the ultimate choice for orthopaedic implants ± a review', Prog Mater Sci, 54, 397±425. Gottlander M, Johansson C B, Wennerberg A, Albrektsson T, Radin S and Ducheyne P (1997), `Bone tissue reactions to an electrophoretically applied calcium phosphate coating', Biomaterials, 18, 551±557. Gupta R and Kumar A (2008), `Bioactive materials for biomedical applications using solgel technology', Biomed Mater, 3, 034005. Harrison N M, Wang X G, Muscat J and Scheffler M (1999), `The influence of soft vibrational modes on our understanding of oxide surface structure', Faraday Discuss, 114, 305±312. Henderson M A (2002), `The interaction of water with solid surfaces: fundamental aspects revisited', Surf Sci Rep, 46, 1±308. Hurlen T and Gulbrandsen E (1994), `Growth of anodic films on valve metals', Electrochim Acta, 39, 2169±2172. Kataoka S, Gurau M C, Albertorio F, Holden M A, Lim S M, Yang R D and Cremer P S (2004), `Investigation of water structure at the TiO2/aqueous interface', Langmuir,

ß Woodhead Publishing Limited, 2011

Metal surface oxidation and surface interactions

139

20, 1662±1666. Kubo T, Orita H and Nozoye H (2007), `Surface structures of rutile TiO2 (011)', J Am Chem Soc, 129, 10474±10478. Labat F, Baranek P and Adamo C (2008), `Structural and electronic properties of selected rutile and anatase TiO2 surfaces: an ab initio investigation', J Chem Theory Comput, 4, 341±352. Langer R and Tirrell D A (2004), `Designing materials for biology and medicine', Nature, 428, 487±492. Larsson C, Thomsen P, Lausmaa J, Rodahl M, Kasemo B and Ericson L E (1994), `Bone response to surface modified titanium implants: studies on electropolished implants with different oxide thicknesses and morphology', Biomaterials, 15, 1062±1074. Lausma J (2001), `Mechanical, thermal, chemical and electrochemical surface treatment of titanium', in Brunette D M, Tengvall P, Textor M and Thomsen P, Titanium in medicine: material science, surface science, engineering, biological responses, and medical applications, Berlin; New York, Springer: 231±266. Lausmaa J, Kasemo B, Mattsson H and Odelius H (1990), `Multi-technique surface charaterization of oxide films on electropolished and anodically oxidized titanium', Appl Surf Sci, 45, 189±200. Lendlein A and Kelch S (2002), `Shape-memory polymers', Angew Chem-Int Edit, 41, 2034±2057. Li P and de Groot K (1993), `Calcium phosphate formation within sol±gel prepared titania in vitro and in vivo', J Biomed Mater Res, 27, 1495±1500. Lindan P J D, Harrison N M, Holender J M, Gillan M J and Payne M C (1996), `The TiO2 (100)(1  3) reconstruction: Insights from ab initio calculations', Surf Sci, 364, 431±438. Lindsay R, Wander A, Ernst A, Montanari B, Thornton G and Harrison N M (2005), `Revisiting the surface structure of TiO2 (110): a quantitative low-energy electron diffraction study', Phys Rev Lett, 94, 246102. Liu X, Chu P K and Ding C (2004), `Surface modification of titanium, titanium alloys, and related materials for biomedical applications', Mat Sci Eng R, 47, 49±121. Long M and Rack H J (1998), `Titanium alloys in total joint replacement ± a materials science perspective', Biomaterials, 19, 1621±1639. Machesky M L, Predota M, Wesolowski D J, Vlcek L, Cummings P T, Rosenqvist J, Ridley M K, Kubicki J D, Bandura A V, Kumar N and Sofo J O (2008), `Surface protonation at the rutile (110) interface: explicit incorporation of solvation structure within the refined MUSIC model framework', Langmuir, 24, 12331±12339. MaÂrcia Soares S, Alex B, GloÂria de Almeida S and Radovan B (2005), `Effect of three distinct treatments of titanium surface on osteoblast attachment, proliferation, and differentiation', Clin Oral Implan Res, 16, 667±675. McGee M A, Howie D W, Costi K, Haynes D R, Wildenauer C I, Pearcy M J and McLean J D (2000), `Implant retrieval studies of the wear and loosening of prosthetic joints: a review', Wear, 241, 158±165. Monti S, Carravetta V, Battocchio C, Iucci G and Polzonetti G (2008), `Peptide/TiO2 surface interaction: A theoretical and experimental study on the structure of adsorbed ALA-GLU and ALA-LYS', Langmuir, 24, 3205±3214. Moskalewicz T, Czyrska-Filemonowicz A and Boccaccini A R (2007), `Microstructure of nanocrystalline TiO2 films produced by electrophoretic deposition on Ti-6Al-7Nb alloy', Surf Coat Tech, 201, 7467±7471. Muscat J and Harrison N M (2000), `The physical and electronic structure of the rutile (001) surface', Surf Sci, 446, 119±127.

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Navarro M, Michiardi A, Castano O and Planell J A (2008), `Biomaterials in orthopaedics', J R Soc Interface, 5, 1137±1158. Niinomi M (2008), `Metallic biomaterials', J Artif Organs, 11, 105±110. Oh S H, Finones R R, Daraio C, Chen L H and Jin S (2005), `Growth of nano-scale hydroxyapatite using chemically treated titanium oxide nanotubes', Biomaterials, 26, 4938±4943. Pilliar R M (1987), `Porous-surfaced metallic implants for orthopaedic applications', J Biomed Mater Res, 21, 1±33. Pilliar R M (1998), `Overview of surface variability of metallic endosseous dental implants: textured and porous surface-structured designs', Implant Dent, 7, 305±314. Raffaini G and Ganazzoli F (2003), `Simulation study of the interaction of some albumin subdomains with a flat graphite surface', Langmuir, 19, 3403±3412. Raffaini G and Ganazzoli F (2004a), `Molecular dynamics simulation of the adsorption of a fibronectin module on a graphite surface', Langmuir, 20, 3371±3378. Raffaini G and Ganazzoli F (2004b), `Surface ordering of proteins adsorbed on graphite', J Phys Chem B, 108, 13850±13854. Raffaini G, Elli S and Ganazzoli F (2006), `Computer simulation of bulk mechanical properties and surface hydration of biomaterials', J Biomed Mater Res A, 77A, 618± 626. Raffaini G and Ganazzoli F (2007), `Understanding the performance of biomaterials through molecular modeling: crossing the bridge between their intrinsic properties and the surface adsorption of proteins', Macromol Biosci, 7, 552±566. Raffaini G and Ganazzoli F (2010), `Surface hydration of polymeric (bio)materials: A molecular dynamics simulation study', J Biomed Mater Res A, 92A, 1382±1391. Richmond G L (2002), `Molecular bonding and interactions at aqueous surfaces as probed by vibrational sum frequency spectroscopy', Chem Rev, 102, 2693±2724. Schwarz F, Wieland M, Schwartz Z, Zhao G, Rupp F, Geis-Gerstorfer J, Schedle A, Broggini N, Bornstein M M, Buser D, Ferguson S J, Becker J, Boyan B D and Cochran D L (2009), `Potential of chemically modified hydrophilic surface characteristics to support tissue integration of titanium dental implants', J Biomed Mater Res B, 88B, 544±557. Sittig C, HaÈhner G, Marti A, Textor M, Spencer N D and Hauert R (1999a), `The implant material, Ti6Al7Nb: surface microstructure, composition and properties', J Mater Sci-Mater M, 10, 191±198. Sittig C, Textor M, Spencer N D, Wieland M and Vallotton P H (1999b), `Surface characterization', J Mater Sci-Mater M, 10, 35±46. Sousa S R, Moradas-Ferreira P, Saramago B, Viseu Melo L and Barbosa M A (2004), `Human serum albumin adsorption on TiO2 from single protein solutions and from plasma', Langmuir, 20, 9745±9754. Sousa S R, Bras M M, Moradas-Ferreira P and Barbosa M A (2007), `Dynamics of fibronectin adsorption on TiO2 surfaces', Langmuir, 23, 7046±7054. Sousa S R, Lamghari M, Sampaio P, Moradas-Ferreira P and Barbosa M A (2008), `Osteoblast adhesion and morphology on TiO2 depends on the competitive preadsorption of albumin and fibronectin', J Biomed Mater Res A, 84A, 281±290. Streicher R M, Weber H, SchoÈn R and Semlitsch M (1991), `New surface modification for Ti-6Al-7Nb alloy: oxygen diffusion hardening (ODH)', Biomaterials, 12, 125± 129. Sun H (1998), `COMPASS: an ab initio force-field optimized for condensed-phase applications ± overview with details on alkane and benzene compounds', J Phys Chem B, 102, 7338±7364.

ß Woodhead Publishing Limited, 2011

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Swamy V, Muscat J, Gale J D and Harrison N M (2002), `Simulation of low index rutile surfaces with a transferable variable-charge Ti-O interatomic potential and comparison with ab initio results', Surf Sci, 504, 115±124. Textor M, Sittig C, Frauchiger V, Tosatti S and Brunette D M (2001), `Properties and biological significance of natural oxide films on titanium and its alloys', in Brunette D M, Tengvall P, Textor M and Thomsen P, Titanium in medicine: material science, surface science, engineering, biological responses, and medical applications, Berlin; New York, Springer: 171±230. Thompson T L and Yates J T (2006), `Surface science studies of the photoactivation of TiO2 ± new photochemical processes', Chem Rev, 106, 4428±4453. Uosaki K, Yano T and Nihonyanagi S (2004), `Interfacial water structure at as-prepared and UV-Induced hydrophilic TiO2 surfaces studied by sum frequency generation spectroscopy and quartz crystal microbalance', J Phys Chem B, 108, 19086±19088. Van der Biest O O and Vandeperre L J (1999), `Electrophoretic deposition of materials', Annu Rev Mater Sci, 29, 327±352. Variola F, Vetrone F, Richert L, Jedrzejowski P, Yi J H, Zalzal S, Clair S, Sarkissian A, Perepichka D F, Wuest J D, Rosei F and Nanci A (2009), `Improving biocompatibitity of implantable metals by nanoscale modification of surfaces: an overview of strategies, fabrication methods, and challenges', Small, 5, 996±1006. VoÈroÈ s J, Wieland M, Laurence R-T, Textor M and Brunette D M (2001), `Characterization of titanium surfaces', in Brunette D M, Tengvall P, Textor M and Thomsen P, Titanium in medicine: material science, surface science, engineering, biological responses, and medical applications, Berlin; New York, Springer: 87±144. Wang R, Hashimoto K, Fujishima A, Chikuni M, Kojima E, Kitamura A, Shimohigoshi M and Watanabe T (1997), `Light-induced amphiphilic surfaces', Nature, 388, 431±432. Wennerberg A and Albrektsson T (2009), `Effects of titanium surface topography on bone integration: a systematic review', Clin Oral Implan Res, 20, 172±184. Yerokhin A L, Nie X, Leyland A, Matthews A and Dowey S J (1999), `Plasma electrolysis for surface engineering', Surf Coat Tech, 122, 73±93. Zhang Z, Fenter P, Cheng L, Sturchio N C, Bedzyk M J, Predota M, Bandura A, Kubicki J D, Lvov S N, Cummings P T, Chialvo A A, Ridley M K, Benezeth P, Anovitz L, Palmer D A, Machesky M L and Wesolowski D J (2004), `Ion adsorption at the rutile±water interface: Linking molecular and macroscopic properties', Langmuir, 20, 4954±4969. Zhitomirsky I (2000), `Electrophoretic hydroxyapatite coatings and fibers', Mater Lett, 42, 262±271. Zhitomirsky I (2002), `Cathodic electrodeposition of ceramic and organoceramic materials. Fundamental aspects', Adv Colloid Interfac, 97, 277±315. Zhitomirsky I and GalOr L (1997), `Electrophoretic deposition of hydroxyapatite', J Mater Sci-Mater M, 8, 213±219.

5.6

Appendix A: Materials and methods for unpublished results

The simulations of the surface hydration were carried out as previously described (Raffaini et al., 2006; Raffaini and Ganazzoli, 2010) using the

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Materials Studio 4.3 simulation package distributed by Accelrys Inc., San Diego CA (2008), available at URL http://www.accelrys.com. using the COMPASS force field (Sun, 1998). The surfaces of the TiO2 polymorphs were generated as a relatively thin slab from the known crystalline coordinates already available in the Materials Studio database by creating a supercell through periodically repeating the crystal cell along the three axes. We then hugely increased the axis perpendicular to the surface of interest, so as to generate a vast empty space that was filled with water molecules at a (local) density of 1 g cmÿ3. Molecular Mechanics (MM) methods were used for the initial geometry optimization of water, and then Molecular Dynamics (MD) runs were performed at a constant temperature (300 K) for 10 ps after an initial equilibration of 0.4 ps. The energy minimizations were carried out up to an energy gradient lower than 4
10ÿ3 kJ Ê ÿ1. The temperature was controlled through the Berendsen thermostat moleÿ1 A and the dynamical equations were integrated with the Verlet algorithm using a time step of 1 fs. The trajectories of the MD runs were then used to calculate the statistical distribution of the water molecules as a function of their distance from the surface.

5.7

Appendix B: Abbreviations and symbols

AFM ASD Ca/P CVD DFT DMTAP DOPC DOPS ELD EPD LCAO LEED MAO MD MM PVD SBF STM SXRD UHMWPE VSFS

Atomic force microscopy Anodic spark deposition Calcium phosphate Chemical vapour deposition Density functional theory Myristic acid derivative of trimethylammonium propane Phosphatidylcholine derivative of the oleic acid Phosphatidylserine derivative of the oleic acid Electrolytic deposition Electrophoretic deposition Linear combination of atomic orbitals Low energy electron diffraction spectroscopy Micro-arc oxidation Molecular dynamics Molecolar mechanics Physical vapour deposition Simulated body fluid Scanning tunneling microscopy Surface X-ray diffraction Ultra high molecular weight polyethylene Vibrational sum frequency spectroscopy

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Surface modification of biomaterials by calcium phosphate deposition J . A . J U H A S Z and S . M . B E S T , University of Cambridge, UK

Abstract: Modification of the surface chemistry of bone replacement materials in order to provide enhanced cell attachment, growth and tissue formation can be achieved via various processes which in turn, are not detrimental to the mechanical properties of the bulk material. The most common involve the deposition of calcium phosphates onto the implant surface which is able to direct cells to an osteogenic phenotype. The most recent techniques of depositing calcium phosphate coatings are discussed in this chapter, illustrating their effects on the mechanical and biological performance of the coated substrates and their potential for future clinical applications. Key words: calcium phosphate deposition, simulated body fluid (SBF), coating(s), apatite, deposition techniques.

6.1

Introduction

The average age of the population is increasing worldwide due to better healthcare and constant improvements in medical research. However, when considering bone and joint replacement surgery due to trauma or disease, there is still the need for better treatments to avoid material failure and the need for revision surgery. Limited supplies of natural bone (autographs and allographs) as well as issues with donor site morbidity and immune rejection indicate the need for artificial alternatives. The inability of bone grafts to replace large, weight-bearing areas of bone indicates metals to be the only suitable substitute. These materials possess sufficient mechanical strength to be placed in loadbearing applications. However, they have serious shortcomings related to osseointegration and the mismatch between their mechanical properties and that of bone, which hinder their long-term stability. The lack of a bioactive bond between metals and natural bone has led to extensive studies relating to surface treatments, most successfully, via the process of plasma spraying [1±3]. The issue of non-osseointegration is not limited to metals. Polymers, which have excellent tuneable biodegradability lack bioactivity and have poor strength and stiffness. However, groups have already begun to illustrate methods of enhancing mechanical properties and inducing bioactivity in certain polymers [4, 5]. These will be discussed in detail later on in this chapter.

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As already mentioned, the absent but essential properties of various biomaterials indicates a requirement for next generation biomaterials. A need for new and improved materials with bone-like mechanical properties, excellent biocompatibility and bioactivity and the ability to degrade and be replaced by natural bone.

6.2

Basic methods and applications

Modification of the surface chemistry of biomaterials in order to provide enhanced cell attachment, growth and tissue formation can be achieved via various processes which in turn, are not detrimental to the mechanical properties of the bulk material. The most common involve the deposition of calcium phosphates onto the implant surface which is able to direct cells to an osteogenic phenotype [6±8]. The typical techniques currently being investigated for metallic and polymer materials are described below. The formation of a chemical bond between human bone and an implant material requires there to be either bone growth onto or into the prosthesis. As far back as in the 1960s, it was realised that calcium phosphate coatings would provide much better fixation of load-bearing implants than using cements such as polymethylmethacrylate (PMMA) [9]. Calcium phosphates such as hydroxyapatite (HA) are suitable since they closely resemble the mineral phase in human bone which is composed of inorganic apatite crystals and organic collagen [10]. Hydroxyapatite, when used in vivo, has shown it is non-inflammatory and does not cause a pathological reaction [11]. Hence, its suitability as a biological coating material for otherwise, `inert' materials [12±18]. Once a calcium phosphate has been coated onto a substrate material, the properties of the coating are required to meet certain criteria set out by the Food and Drug Administration USA (FDA) and International Standard Organization (ISO) [19]. It calls for the minimum crystallinity to be 62% and purity of the coating to be 95% with the tensile strength and shear strength being less than 50.8 and 22 MPa, respectively. All commercially produced coatings meet these standards and any currently investigated CaP deposition techniques work towards attaining these values in order to provide improved bioactive fixation of implants which would aid the long- and short-term success of implants. Typical calcium phosphate coating techniques include ion beam assisted deposition, plasma spray deposition, magnetron sputtering, and non-thermal biomimetic methods performed under normal atmospheric conditions. Plasma spraying techniques are the main route used commercially to coat metals such as titanium with HA. It is a high temperature process which is more complex and expensive than other techniques such as biomimetics. This technique has been used clinically for many years and in a recent report, demonstrated its success in patients after 13 years implantation [20]. In this chapter more attention has been directed towards the recent research

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techniques being investigated as the potential alternatives to plasma spraying which is now, well established and used for commercial applications.

6.2.1

Thermal spray deposition

Plasma spraying was established as the most widely used commercial method of preparing CaP coatings in the 1980s. Since then, numerous research groups and companies have used this technique to coat biomaterials such as Ti and Co-Cr alloys due to being a reproducible and cost-effective technique [21±30]. The spraying process involves injecting a HA powder into a plasma flame at high temperature. The powder is heated and forced at high pressure and velocity, towards the substrate material (Fig. 6.1). A molten particle of HA can have an outside temperature of at least 1000 ëC. Owing to this very high temperature, the thin outer layer of each HA particle will inevitably undergo phase transitions. This surface should be sufficiently large to be able to plasticise the outer layer and allow a dense and strongly adhesive coating to be formed. It must also be small enough not to affect the overall crystalline phase deposited on the substrate material [31±33]. Factors in plasma spray deposition that can affect the phase purity of the resultant coating have been noted to be related to the type of gas that is used and the rate at which it accelerates the powder towards the substrate. The type of plasma spraying performed can be abbreviated according to the atmosphere surrounding the process. If air is used, it is denoted as APS (air plasma spraying). This involves the passing of gases or gas mixtures through a flame. An electric arc causes the gas to dissociate and form ions which collide with atoms in the gas which forms a plasma flame. CaP powder is passed into the plasma flame in which it is melted and then accelerated towards the substrate. Amongst the specific atmospheres available for plasma spraying the most commonly used at present is VPS (vacuum plasma spraying) or LPPS (low pressure plasma spraying). This is able to improve the quality of the coating, more crystalline that that produced by APS [34], as well as its adhesive strength. However, this comes at a price since the technique is more costly compared with conventional techniques. The type of coating material and substrate used can also affect the outcome of the process as can the temperature of both materials. All these and other parameters as mentioned below, need to be considered when performing plasma spraying of CaP [22, 35±37]. Performing VPS, i.e. at lower pressures, allows for an increase in the velocity of particles and for any undesirable reaction products to be minimised, products such as amorphous CaP or calcium oxide [38, 39]. The main disadvantage with this technique is that the substrate surface temperature will be increased, but this is counteracted by the resultant coating which has good adhesion and high density [40].

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6.1 Plasma-sprayed HA coating on titanium alloy substrate (a) and human osteoblast-like cell (HOB) growth on HA-coated titanium alloy for 3 days (b) [170].

High-velocity oxy-fuel (HVOF) spray deposition was first developed in the 1980s and involves the mixture of a gaseous or liquid fuel and oxygen; a combustion spray technique. The combustion process generates heat and accelerates particles at very high velocities towards the substrate surface. It ensures a coating with good adhesive strength. The temperature of the flame is lower than that used in conventional plasma spraying, minimising over-heating

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or vaporisation of the particles [40]. Once the particles adhere to the substrate, their heat is quickly dissipated to the substrate, which allows for rapid solidification of the coating, helping to retain its crystallinity, and an improved bonding strength compared with other techniques [41±43]. However, the resultant coating has reduced density due to insufficient melting of each particle surface. The processing conditions, as for plasma spraying, have a significant effect on the structure of the resultant coating and its mechanical and in vitro properties [44]. Using HA as a coating in composite-form, with the addition of YSZ (yttrium-stabilised zirconia) and titania, improved the mechanical performance of the coatings, but the processing technique caused chemical reactions between the individual components which had a detrimental effect on the performance of the coated material in vitro [45, 46]. Animal studies have shown that plasma sprayed HA-coated implants allow faster bone ingrowth with a higher percentage of bone contact on the samples plasma sprayed with HA [47, 48]. As well as these in vivo results, there has been much clinical success with HA plasma-sprayed metals [49]. The most recent report, mentioned earlier, showed how well implants had been accepted in patients after 13 years implantation [20]. However, it has also been documented that these uncemented prostheses are difficult to extract when revision surgery is required due to ongoing pain or infection [9]. There have also be problems in a few implants with third-body wear due to coatings containing high amounts of amorphous CaP and being worn away causing inflammatory problems and the need for revision surgery [50, 51].

6.2.2

Ion beam assisted deposition

This deposition methodology comprises an electron beam which vaporises a hydroxyapatite target, while an argon ion beam is focused onto the substrate to assist deposition. It creates a deposit of HA with high adhesive strength; the strength of which increases with increased current. This increase in bonding strength is due to the atomic interaction at the interface between substrate and HA coating. Ion beam deposition also allows for densification of the coating and lower thermal stresses to exist, preventing delamination of the coating. All deposited layers that are formed by this technique are amorphous, regardless of the current used, which indicates that the coatings need to be heat-treated in order to provide sufficient stability to the coatings in vivo. The final level of crystallinity, and hence, degree of bioactivity, is also important and is dependent on the time, temperature and amount of water vapour present during the deposition process [52, 53]. A graded crystallinity can be achieved by heating the substrate during the deposition process, which can be beneficial to the rate of bone growth and mechanical stability and strength of the coating. It has been proprosed that the presence of an amorphous layer at the top of such a coating can allow for faster osteointegration once placed in vivo [54].

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Pulsed laser physical vapour deposition

Physical vapour deposition is a complex process which involves the laser ablation of a target material and the deposition via nucleation and film growth of the ablated material onto a substrate. The individual steps are vital for the final film crystallinity and stability. Pulsed laser deposition can also be used to create multicomponent stoichiometric films from one target [55]. Pulsed laser can be used to synthesise metastable materials which would otherwise with standard techniques, be difficult to fabricate. Nano-crystalline coatings can also be created as well as composite coatings [55±57]. Numerous calcium phosphates have been used as coating materials including HA, octacalcium phosphate (Ca8H2(PO4)6.5H2O (OCP)) and alpha and beta-tricalcium phosphates [58±75]. Pulsed laser deposition was used by various groups to form thin HA coatings on polished substrates [58±61]. The process involves ablation of a HA target in a water vapour atmosphere and depositing the ablated HA material onto a heated substrate (400±800 ëC). Nelea et al. [58, 59] grew thin films of HA onto Ti alloy substrates which had been pre-coated with a buffer layer of TiN. This buffer layer improved the adhesion of the HA film to the substrate. The benefits of this TiN layer were reinforced when Ti alloys were not pre-treated and were allowed to form an oxide layer between the film and the substrate [65, 66]. The presence of this oxide layer was a point of weakness and was detrimental to both the strength and microhardness of the films. Analysis of the coating revealed they had a rough surface which would be advantageous for osteoblast cell adhesion. As well as the effect of surface roughness, the beneficial effects of increased crystallinity and laser fluence were demonstrated when human osteoblast cells were cultured on thin HA films created by pulsed laser physical vapour deposition [62]. Enhanced cell proliferated occurred on these films due in part, to the improved HA coating crystallinity which would result in reduced dissolution and release of calcium and phosphorus ions. Nanosized octacalcium phosphate (OCP) deposited onto titanium has demonstrated good bioactivity and is more resorbable than HA [63]. Both fibro- and osteoblast cells were able to differentiate and grow on these surfaces and adhered strongly to the OCP films formed by physical vapour deposition.

6.2.4

Micro-arc oxidation

Micro-arc oxidation (MAO), also known as plasma electrolytic oxidation, anodic spark deposition or micro-arc discharge oxidation, is an electrochemical surface treatment which results in the generation of oxide layers on metals. MAO can be performed at ambient temperature and used on substrates with complex geometries. It has been used to form various calcium phosphate coatings on metal alloys [76±80].

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Altering the processing parameters can determine the morphology and composition of the calcium phosphate coating formed. Using various applied voltages, for example changing the cathode voltage during coating synthesis will affect both the phase composition and morphology of the resultant film [76, 81]. However, an optimum needs to be found since voltages that are too high, above a few hundred, will destroy the coating, and voltages that are too low will lead to pore formation which will be detrimental to the strength of the coatings. As well as varying equipment parameters, the concentration of the HA used in Ti-based coatings on titanium alloys can also be adjusted [78]. By increasing the HA content, the behaviour of these coatings in simulated body fluid (SBF) will change from anatase (low crystallinity) and amorphous CaP to the presence of rutile. As the HA content is further increased, the crystallinity of the coatings is improved, elevating their bioactivity. Creating composite coatings with a HA filler has demonstrated suitable bioactivity and the ability to control the calcium phosphate phase composition and mechanical properties of these coatings when produced by MAO. Ion-substituted HA coatings, for example strontium-HA [82], made via MAO methods has also proven to offer promising results in vitro.

6.2.5

Magnetron sputtering deposition

Magnetron sputtering is a high-rate vacuum coating technique that allows the deposition of many types of materials, including metals and ceramics, onto as many types of substrate materials by the use of a specially formed magnetic field applied to a diode sputtering target. It allows a faster deposition rate at lower pressures compared with other techniques and is able to create strongly adhesive coatings on complex geometries including those made of heat-sensitive substrates such as polymers. Due to the desirable properties of this technique, many researchers have used this process to apply CaP films onto metallic and non-metallic implants [83±117, 170]. It has demonstrated to be a promising method for forming a biocompatible coating on metal, plastic and ceramic substrates since a wide variety of materials can be used and the processing parameters and postprocedure heat-treatments such as post deposition or in situ annealing can be applied [104±111] (Plate II between pages 208 and 209). Altering parameters such as the gas pressure (argon) can have a marked effect on the HA coating crystallinity, causing dehydroxylation and amorphous calcium phosphate formation when the gas pressures are too high [107]. In the mid-1990s, Wolke et al. [83] were among the first to try radiofrequency magnetron sputtering as an alternative method to deposit thin films of HA onto titanium substrates. A crystalline layer was formed on the titanium surface with a preferred crystallographic orientation (001) and a uniform and dense structure. Developments in magnetron sputtering led to findings that this technique could be used to vary the coating thickness which could subsequently

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be heat-treated to form a crystalline structure. Such heat-treated coatings have high adhesive strength, but due to the annealing process, are prone to cracking when thicker coatings are applied (>0.1 microns). Despite these drawbacks, magnetron-sputtered HA coatings allow for a strong bond to be formed with adjacent bone when implanted via the growth of fibrillar collagen matrices and carbonate apatite crystals. The microstructure and bonding strength of HA coatings produced using magnetron sputtering on Ti substrates can be improved by the formation of a TiN buffer layer and by altering the substrate temperature during the sputtering process. Using the TiN buffer provides an improved bonding strength between the substrate and HA film and heating the substrate allows the coating to be transformed from an amorphous to a more crystalline HA structure [75]. As well as metals, polymers and ceramics can also be coated successfully using magnetron sputtering producing fully dense and strongly adherent calcium phosphate coatings [75, 87, 91±99]. For such substrates, ion-substituted HAs and composite coatings of HA/TiO2 can be applied [75, 87, 108]. The use of substituted apatites such as silicon-substituted HA has been shown to encourage bone-bonding much more than pure HA and has been used as a coating material for various load-bearing implant materials [100, 101].

6.2.6

Electrophoretic deposition

Electrophoretic deposition (EPD) has been used to form CaP coatings on various substrate materials. This process involves an electric field which is employed to deposit charged particles (be that cathodic or anodic) in a liquid (not water) onto a conducting substrate of opposite charge. It is a highly versatile and low cost procedure that allows any 3D object (including porous materials) to be uniformly coated, which has been used successfully to deposit bioactive coatings onto medical implants. Varied thicknesses of calcium phosphate coating can be formed on metal substrates by changing the electric field and deposition time during electrophoresis. Resultant coatings are uniformly distributed and densified, but due to differences in thermal expansion coefficients of the metal substrate and ceramic coating can lead to cracking. By performing repeat depositions at room temperature and by varying the dispersant type (e.g. glycol for ethanol), the issue of coating cracking can be minimised [118, 119]. Another process applied to improve the adhesion between substrate and CaP coating is the use of an intermediate layer. Silica, calcium-silica or titania can be deposited onto metals allowing the `second', HA coating to bind strongly and without the formation of any cracks. The use of nanoparticulate HA coatings alone and as this `second' layer has indicated a new trend in electrophoretic techniques for biometals which also overcome the issue of coating cracking [120±122]. The intermediate layer acts as a diffusion barrier that prevents

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decomposition of the nano-sized HA which would otherwise occur due to ion migration from the metal substrate into the HA. The intermediate coatings are also an effective chemical barrier to the release of potentially inflammatory metal ions from the bulk implant material. Recently, composites have also gained interest as materials for coatings applied by electrophoresis. Coatings of silicon-substituted HA and polycaprolactone (PCL, a known biodegradable polymer) on titanium substrates have shown improved bonding strength compared to Si-HA alone due to the addition of PCL. This polymer can degrade over time and does not hinder the bioactivity of the HA [123]. Substrates besides titanium and titanium alloys, such as stainless steel, have also been coated effectively with calcium phosphates and calcium phosphate composites using electrophoretic techniques [124±126]. Composite HA-chitosan coatings on stainless steel demonstrated that a uniform and well-adhered coating could be produced with HA nanoparticle preferred orientation being increased with increasing chitosan content.

6.2.7

Electrochemical deposition

The process of electrochemical deposition involves the decomposition of an aqueous electrolyte which, for the formation of an apatite coating on biomaterials, must contain calcium and phosphorus ions. Bioactive coatings obtained by electrochemical deposition are composed of interlocking calcium phosphate crystal networks. These are formed via a nucleation and growth process which is affected by various factors such as electrolytic solution composition (additions of NaNO3 or NaF) or concentration, electric current density, deposition time, and solution pH. These factors can be tailored to control the purity, crystallinity, stoichiometry, morphology and mechanical strength of the resulting coatings. Varying the current density has a direct effect on the local pH of the solution and can determine the morphology and crystalline structure of the resultant coatings [127]. The electrolyte temperature can affect the amount of coating that is deposited on the surface of metals such that when the temperature is increased, more calcium and phosphate ions are absorbed. The degree of aqueous solution supersaturation used to form CaP coatings can also control the morphology and length of apatite crystals [128]. All these parameters can aid the bioactivity of the final coated metal, enhancing the rate at which early fixation occurs. Numerous modified electrochemical techniques have been used to improve the adhesive strength between apatite coatings and substrates. These have included the formation of a composite coating by the combination of HA with vinyl acetate and the anodic oxidation of TiO2 nanotubes onto titanium substrate prior to HA deposition [129±131]. Vinyl acetate alters the morphology of the HA crystals and causes a significant increase in the adhesive strength of the

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coating. The anodic pre-treatment of the metal substrate leads to the formation of a TiO2/dicalcium phosphate dihydrate (DCPD) coating which can be converted to a more crystalline structure by immersion in an alkaline solution. Simulated body fluid is a solution that can be used as the electrolyte source during electrochemical CaP deposition. SBF at different temperatures can affect the phase composition of the coating produced. With increasing temperatures, the structure of the CaP coating becomes more crystalline and carbonate-rich. Using SBF during electrochemical deposition at body temperature and pH (around 37 ëC and pH 7.40) causes the calcium phosphate crystals to be thin and nano-sized. Once these crystals are formed, spheres of water-rich amorphous CaP can grow on this layer which, depending on the current density used, will be transformed into crystalline needles of HA. The higher the current density, the more crystalline the coating and, hence, the ability to control the solubility and adhesive strength of these coatings [132]. Electrochemically deposited CaP coatings can be formed with ionic substitutions in their structure, such as carbonate and fluoride. Carbonatesubstituted apatite can be coated onto titanium plates using electrolytes containing calcium and phosphate ions (SBF levels) but without the use of a pH buffer. These can sustain osteoblast cell activity [133]. The addition of fluoride ions to apatite can be applied to both orthopaedic and dental percutaneous implants, retaining the bioactivity of the HA and passing on the beneficial antibacterial properties of fluorine. Along with these forms of substituted apatites, composites such as CaPs with chitosan and collagen, HA with poly(vinyl acetate) (PVA) and polypyrrole are also proving to give excellent results, both biologically and mechanically, even on NiTi shape-memory orthopaedic alloys [134±137]. The addition of polypyrrole to apatite during the electrochemical deposition using SBF as the electrolyte promotes the incorporation of carbonate ions into the calcium phosphate coating. The carbonate-substituted HA enhancing the bioactivity of the coated material.

6.2.8

Sol-gel methods

Sol-gel processing is a wet chemistry approach which results in ceramic coatings with an exact chemical and microstructural composition. The desired coating content is mixed at low temperature in solution as a colloidal suspension of inorganic particles and then heat-treated at high temperatures (up to 1000 ëC) in order to densify the coating on the substrate. Sol-gel processing allows for the production of thin films with excellent adhesion properties which can be applied to complex geometries and to non-metallics such as polymers due to the low sintering temperatures which can be used to heat-treat the `green' coatings. Significant amounts of research have reported the benefits of using this technique to deposit calcium phosphate onto metallic substrates. Following the sol-gel treatment, coated samples are heat-treated which affects the stability,

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purity and crystallinity of the resultant calcium phosphates. Trace amounts of calcium oxide and beta-tricalcium phosphate (beta-TCP) have been noted when heat-treating HA [138, 139], but these are undesirable phases due to beta-TCP being much more soluble than HA (reducing bioactivity and coating adhesive strength) and CaO lacking biocompatibility and dissolving in body fluids at a faster rate than TCP. However, these phases can be avoided and monophasic HA produced by increasing the ageing time to more than 24 hours. Variations in the processing parameters can also control the quality of the resulting coating material. For example, by increasing the calcining temperature and changing the pH of the sol solutions a HA coating with high purity and crystallinity can be produced, which in turn, will have better coating stability in vivo and induce optimum levels of bioactivity. As well as controlling the processing parameters, starting material compositions can also be tailored. Substituted forms of HA such as fluoride- and silicon-HA are beneficial at increasing the bioactivity of the calcium phosphate coatings, and increasing the crystallinity of the apatite film. Pre-treating substrate surfaces with titanium oxide or calcium phosphate compounds using electrophoresis promote both the mechanical strength of the apatite coatings and their bioactivity [140, 141]. A similar favourable result in bonding strength is witnessed when magnetron sputtering and electrochemical deposition techniques employ TiO2 pre-coating [108, 130].

6.2.9

Hot isostatic pressing

The aim of hot isostatic pressing (HIP) is to overcome the problems associated with plasma spraying of HA such as coating porosity and weak bonding to the substrate. Prior to HIP being performed, the HA powder is applied to the surface of the substrates by techniques such as mixing with water and then air-spraying onto Ti substrates. At temperatures above 800 ëC the Ti coated samples are coldpressed with a pressure of around 700 bar to produce optimum, well-adhered coatings. This ensures coatings which have greater density than with plasma spraying alone. In vivo testing illustrated that using higher pressures led to improved bone/implant bonding values than sand-blasted Ti implants or those cold-pressed at lower pressures [142]. Despite the success in vivo, it was found that this technique contaminated the coatings with metal and SiO2 particles due to the glass tubing that was used during the HIP process. More recent work has involved plasma spraying Ti alloy substrates (the most successful commercial form of HA coating to-date) and then performing hot isostatic pressing on the resultant materials [143]. The HIP process allows for a reduction in the porosity of the coatings since pores, especially those of 10± 300 nm are detrimental to the mechanical strength of coatings. The HIP process compresses the pores and dramatically increases the density of the coatings, hence enhancing the mechanical properties of the HA.

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Hydrothermal forms of the HIP process are also becoming a more prominent method of depositing dense layers of HA onto substrates. It is a double layer technique which allows for the hydrothermal treatment to be performed during the HIP, creating coatings on curved surfaces with good bonding strength [144, 145].

6.2.10 Biomimetic coatings Calcium phosphate which is found in natural hard tissues can be produced spontaneously in physiological, supersaturated solutions at low temperatures. It can be used to grow bone-like apatite on potential implant materials and is particularly suitable for the coating of biodegradable polymeric materials and degradable tissue engineering scaffolds [146±153]. The first group to grow calcium phosphate on a substrate using an in vitro biomimetic process was Kokubo et al. in 1990 [147, 154]. They demonstrated the ability of this solution, with ion concentrations close to those of human blood plasma, to induce apatite crystal formation on CaO±SiO2-based glasses. The thickness of this calcium phosphate coating can be controlled with most bioactive samples forming an amorphous or amorphous-crystalline structure on their surface. Since its development, simulated body fluid (SBF) has become one of the simplest in vitro techniques for observing the potential in vivo bioactivity of a material. When using SBF as the basic physiological solutions for apatite formation, modifications in experimental conditions such as temperature, time or ion concentration allow for the tailoring of the apatite formed and the creation of biomimetic coatings on otherwise, `inert' materials. These bioactive coatings cause the formation of a direct chemical bond with bone when placed in the body. Increasing the concentration of SBF to twice that of normal amounts leads to apatite layers forming on polymers such as polyethylene terephthalate. This occurs due to an increased ionic activity but the formation of apatite with Ca/P ratios which are lower and more bioresorbable. Bone-like apatite can be formed on polymers and metals by increasing the concentration of SBF. However, reducing the Mg2+ and HCO32ÿ ion content in this solution allows for faster apatite nucleation and growth due to the removal of the known inhibitory effect of these ions. As well as increasing the concentration of SBF and removing known inhibitory ions, the addition of nucleating agents to the solution can also induce apatite precipitation on the surface of materials. Chemical or physical modification of the substrate surface can also be beneficial to apatite formation. Certain negatively charged surface ±OH groups are able to promote CaP formation, including Si±OH, Ti±OH and Zr±OH which along with being anionic, have a certain spatial arrangement favourable for apatite formation. When considering polymers, surface groups such as ±COOH, ±CONH2, ±OH and ±NH2, have an

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apatite nucleating effect, whereas ±CH3 surface groups are unable to induce such an effect [146, 155]. The influence of a moisture-insensitive, corrosion technique such as sodium hydroxide treatment followed by a heat-treatment can allow the formation of an apatite layer on metals such as Ti and its alloys. A sodium titanate mesh-like surface structure is formed which allows a CaP layer to develop in vivo. Non-surface damaging processes, i.e. without the need for corrosion or surface etching, can be performed which are not detrimental to the mechanical properties of the substrate. Repeated dipping of substrates into a supersaturated calcium phosphate solution and removing and air-drying allows the growth of calcium phosphate crystals on substrate materials. The mechanism of apatite growth is thought to be via an evaporation-induced surface crystallisation process. However, the mechanical stability and bioactivity of these types of coatings need to be analysed. A similar dipping technique to form a nanoCaP rich-hydrogel composite coating on substrates has demonstrated good adhesive strength and promising results when tested in vitro in SBF (Fig. 6.2) and with human osteoblast-like cells [156]. Inducing bioactivity in polymers such as polycaprolactone (PCL), polyethylene terephthalate (PET) and polylactic acid (PLA) can be achieved by biomimetic processes which use bioactive glass particles to stimulate apatite nuclei formation on the substrates in SBF or by alkaline treatments which partially hydrolyse the surface of polymers and generate oxygen-containing functional groups for calcium and phosphate ion interaction. An apatite layer is able to form on these biodegradable polymers after only 24 hours in SBF. The apatite layer formed via the alkali treatment produces coatings with adhesive strengths which make them highly suitable for scaffold applications in bone tissue engineering, outperforming composites currently used clinically for bone replacement. Starch-based biodegradable polymers have gained much interested for use as bone replacement materials with bioactive coatings. Surface modification treatments such as UV radiation, the use of ethylene oxide or the combination of these polymers with sodium silicate gels can allow for apatite nucleation sites to be formed, which is particularly beneficial for porous polymer scaffolds [157, 158]. As well as the success of coating biodegradable polymers with apatite, composite scaffolds of degradable polymers and ceramics have recently been enhanced by the use of biomimetics to form a bioactive coating. Although ceramic-reinforced polymer scaffolds are an improvement on conventional ceramic substitutes, the osteoinductive ability of these scaffolds must be further enhanced to make them suitable for bone tissue engineering. Apatite-coated porous PLA/nanoHA is one such composite which has benefited from being biomimetically coated with apatite resulting in significantly increased cell proliferation and growth rates compared with un-coated HA/PLA composites [159].

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6.2 Apatite formation on a nano-hydroxyapatite±hydrogel coating after 14 days in simulated body fluid at low (a) and high (b) magnifications.

6.3

Strengths and weaknesses

The most commercially used technique for producing HA coatings on orthopaedic and dental implants is plasma spraying, which has been successfully used for many years. However, this process suffers from drawbacks such as poor adhesion properties, non-uniformity and cracking. Due to the high temperatures that are required for this process, it is impossible to add biologically active molecules to the coatings during the apatite layer preparation. One alternative technique being studied is micro-arc oxidation (MAO). This is a process that can be carried out at room temperature for components with

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complex geometries. It is a simple, economical technique for producing calcium phosphate coatings with different surface textures on metal substrates. The production of a porous coating using the MAO technique can enhance the anchorage of the implant to the adjacent bone tissue and can allow the incorporation of antibiotics to reduce infect and implant rejection. Comparisons between electrochemical and biomimetic deposition techniques have shown that the former produces an apatite at a faster rate, but that these coatings typically contain impurities and need post-deposition treatment. The biomimetically deposited CaP layers are much thicker and uniform, especially when coating porous substrates which, with plasma techniques as well as electrochemical treatments, is difficult to achieve. However, there exist differences in morphology which lead to several forms of apatite crystal growth; the edges of samples inhibiting apatite growth due to increased surface energy. Biomimetically formed coatings are preferred by cells due to better proliferation and growth when compared with electrochemical deposition techniques [133]. The success and continuing development of biomimetic techniques are due to its ability to form apatite on all surface topographies and geometries. It is able to incorporate biomolecules and drugs which can promote cell adhesion and growth and can help to miminise the problems associated with implant rejection such as infection. Alternative biomimetic methods have also grown out of the standard techniques mentioned earlier in this chapter. Routes such as layer-bylayer assembly and self-assembled monolayers (SAMs) to incorporate biomolecules such as proteins are all at promising, developmental stages of research [160, 161].

6.4

Future trends

Many techniques which have been used for the deposition of calcium phosphate onto metals or polymers are now being further developed for the use of drug and protein delivery. As well as providing biological anchorage of implants by the formation of an apatite layer, research is being directed towards the addition of bactericidals, glycoproteins and antibiotics to prevent implant rejection and to further accelerate the process of apatite formation. The process of magnetron sputter deposition has been used to create coatings of HA with additions of the known antibacterial agent, silver [162]. Using this technique to add silver to a HA coating allows for a significant reduction in the number of colonies of both S. epidermidis and S. aureus, both of which are bacteria commonly found at surgical sites following orthopaedic implantation. Loading CaP coatings with antibiotics such as amoxicillin and gentamicin is also beneficial, providing implant sites with sustained levels of drug release at therapeutic doses to prevent infection following surgery [163]. The process of electrophoretic deposition has recently made advances allowing the use of carbon nanotubes (CNTs) and heparin-doped organic-

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inorganic composites as coatings for medical applications. The CNTs form a homogeneous surface coating on polymer foams providing a nano-structured surface topography on which apatite crystals are able to nucleate and grow in SBF. The CNTs present more sites for apatite nucleation which accelerate the precipitation of these bioactive crystals, increasing the bioactivity of polymer scaffolds [164]. The use of SBF for biomimetic coating formation has become extremely common in the last 15±20 years. Whether the process involves the use of SBF with ion concentrations similar to human blood plasma or varying its ion content and concentration to induce apatite formation, it is a solution that was primarily set up to allow monitoring of the potential bioactivity of materials. Analysis of the differences between SBF and human blood plasma (i.e. SBF with proteins) may allow for a better understanding of the processes involved during apatite formation [165] and for further improvements in the apatite coatings produced in vitro to be achieved. The deposition techniques discussed in this chapter are not only successfully being applied to create CaP coatings on biomaterials, but are being developed to deposit coatings such as TiN with the addition of silver to form antibacterial coatings for medical applications [166]. These are biocompatible coatings which can provide an implant with reduced risk of infection and improved boneimplant stability. Along with the techniques mentioned here, research groups are continuing to evaluate new approaches to depositing calcium phosphates with the same success as those already used commercially, such as plasma spraying. A technique known as ultrasonic spray pyrolysis (USP) is currently being investigated which is a simple and less costly method of depositing apatite coatings [167± 169] and which has already demonstrated the ability to create homogeneous, highly pure and crystalline HA coatings in air at ambient temperatures and with good adhesive properties.

6.5

Sources of further information and advice

· Thin Calcium Phosphate Coatings for Medical Implants, Eds B Leon and J A Jansen, Springer Science, 1st edn, 2009. · Fifteen Years of Clinical Experience with Hydroxyapatite Coatings in Joint Arthroplasty, Eds J-A Epinette, M T Manley, Springer-Verlag France, 2004. · Characterization and performance of calcium phosphate coatings for implants, Eds E Horowitz, J E Parr, ASTM Committee F-4 on Medical and Surgical Materials and Devices, Philadelphia, PA, 1994. · http://www.msm.cam.ac.uk/ccmm/ · http://www.plasma-biotal.com/ · http://www.isotis.com/ · http://www.nanowerk.com/nanotechnology/nanomaterial/

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References

[1] Willmann G (1999), `Coating of implants with hydroxyapatite material connections between bone and metal', Adv Eng Mater, 1, 95±105. [2] de Groot K, Geesink R G T, Klein C P A T, Serekian P (1987), `Plasma-sprayed coatings of hydroxylapatite', J Biomed Mater Res, 21, 1375±1381. [3] Cook S D, Thomas K A, Kay J F, Jarcho M (1988), `Hydroxylapatite coated titanium for orthopaedic implant applications', Clin Orthop Rel Res, 232, 225±243. [4] Gomes M E, Reis R L, Cunha A M, Blitterswijk C A, de Bruijn J D (2001), `Cytocompatibilty and response of osteoblastic-like cells to starch-based polymers: effect of several additives and processing conditions', Biomater, 22, 1911±1917. [5] Auclair-Daigle C, Bureau M N, Legoux J-G, Yahia L'H (2005), `Bioactive hydroxyapatite coatings on polymer composites for orthopaedic implants', J Biomed Mater Res, 73A, 398±408. [6] Oliveira A L, Malafaya P B, Reis R L (2003), `Sodium silicate gel as a precursor for the in vitro nucleation and growth of a bone-like apatite coating in compact and porous polymeric structures', Biomater, 24, 2575±2584. [7] ter Brugge P J, Wolke J G C, Jansen J A (2002), `Effect of calcium phosphate coating crystallinity and implant surface roughness on differentiation of rat bone marrow cells', J Biomed Mater Res, 60, 70±78. [8] Chang Y, Stanford C M, Wefel J S, Keller J C (1999), `Osteoblastic cell attachment to hydroxyapatite-coated implant surfaces in vitro', Int J Oral Max Impl, 14, 239±247. [9] Furlong R J, Osborn J F (1991), `Fixation of hip prostheses by hydroxyapatite ceramic coatings', J Bone Jt Surg, 73B, 741±745. [10] Mann S (2001), Biomineralization: principles and concepts in bioinorganic materials chemistry, 1st edn, Oxford, Oxford University Press. [11] Soballe K, (1993) `Hydroxyapatite ceramic coating for bone implant fixation: Mechanical and histological studies in dogs', Acta Orthop Scand, 64 Suppl, 1±58. [12] Shi D, Wen X (2006), `Bioactive ceramics: structure, synthesis, and mechanical properties', in Introduction to biomaterials, Ed Shi D, Beijing, Tsinghua University Press, World Scientific, 13. [13] Hench L L (1998), `Bioceramics', J Am Ceram Soc, 81, 1705±1728. [14] Suchanek W, Yoshimura M (1998), `Processing and properties of hydroxyapatitebased biomaterials for use as hard tissue replacement implants', J Mater Res, 13, 94±117. [15] Ong J L, Chan D C N (2000), `Hydroxyapatite and their use as coatings in dental implants: a review', Crit Rev Biomed Eng, 28, 667A±707A. [16] Zhao B H, Lee I S, Bai W, Cui F Z, Feng H L (2005), `Improvement of fibroblast adherence to titanium surface by calcium phosphate coating formed with IBAD', Surf Coat Technol, 193, 366±371. [17] Luo Z S, Cui F Z, Feng Q L, Li H D, Zhu X D, Spector M (2000), `In vitro and in vivo evaluation of degradability of hydroxy apatite coatings synthesized by ion beam-assisted deposition', Surf Coat Technol, 131, 192±195. [18] Xue W C, Tao S Y, Liu X Y, Zheng X B, Ding C X (2004), `In vivo evaluation of plasma sprayed hydroxyapatite coatings having different crystallinity', Biomater, 25, 415±421. [19] FDA (1992), Calcium phosphate (Ca-P) coating draft guidance for preparation of FDA submissions for orthopedic and dental endosseous implants. Washington, DC, Food and Drug Administration, pp. 1±14.

ß Woodhead Publishing Limited, 2011

160

Surface modification of biomaterials

[20] Slack R, Tindall A, Shetty A A (2006), '15-year follow-up results of the hydroxyapatite ceramic-coated femoral stem', J Ortho Surg, 14, 151±154. [21] Ji H X, Ponton C B, Marquis P M (1992), `Microstructural characterization of hydroxyapatite coating on titanium', J Mater Sci ± Mater M, 3, 283±287. [22] Tsui Y C, Doyle C, Clyne T W (1998), `Plasma sprayed hydroxyapatite coatings on titanium substrates Part 2: optimisation of coating properties', Biomater, 19, 2031±2043. [23] Inagaki M, Kameyama T (2007), `Phase transformation of plasma-sprayed hydroxyapatite coating with preferred crystalline orientation', Biomater, 28, 2923±2931. [24] Kohn D H (1998), `Metals in medical applications', Curr Opin Solid St M, 3, 309± 316. [25] Weng J, Liu Q, Wolke J G C, Zhang X D, de Groot K (1997), `Formation and characteristics of the apatite layer on plasma-sprayed hydroxyapatite coatings in simulated body fluid', Biomater, 18, 1027±1035. [26] Zheng X B, Huang M H, Ding C X (2000), `Bond strength of plasma sprayed hydroxyapatite/Ti composite coatings', Biomater, 21, 841±849. [27] Gledhill H C, Turner I G, Doyle C (2001), `In vitro dissolution behaviour of two morphologically different thermally sprayed hydroxyapatite coatings', Biomater, 22, 695±700. [28] Dong Z L, Khor K A, Quek C H, White T J, Cheang P (2003), `TEM and STEM analysis on heat-treated and in vitro plasma-sprayed hydroxyapatite/Ti-6Al-4V composite coatings', Biomater, 24, 97±105. [29] Khor K A, Dong Z L, Quek C H, Cheang P (2000), `Microstructure investigation of plasma sprayed HA/Ti6Al4V composites by TEM', Mater Sci Eng A, 281, 221± 228. [30] Chou B Y, Chang E (1999), `Microstructural characterization of plasma-sprayed hydroxyapatite-10 wt% ZrO2 composite coating on titanium', Biomater, 20, 1823± 1832. [31] Cook S D, Thomas K A, Kay J F, Jarcho M (1988), `Hydroxylapatite coated titanium for orthopaedic implant applications', Clin Orthop Rel Res, 232, 225±243. [32] Mayer Y, Collier J B, Hanes C K (1986), `Enhanced early fixation of porous coated implants using tricalcium phosphate', Trans 32nd Annual Meeting Orthop Res Soc, 348. [33] Wolke J G C, de Blieck-Hogervorst J M A, Dhert W J A, Klein C P A T, de Groot K (1992), `Studies on thermal spraying of apatite bioceramics', J Thermal Spray Tech, 1, 75±82. [34] Gross K A, Berndt C C, Herman H (1998), `Amorphous phase formation in plasma-sprayed hydroxyapatite coatings', J Biomed Mater Res, 39 407±414. [35] Lugscheider E, Barimani C, Eckert P, Eritt U (1996), `Modeling of the APS plasma spray process', Comp Mater Sci, 7, 109±114. [36] Gross K A, Berndt C C (1998), `Thermal processing of hydroxyapatite for coating production', J Biomed Mater Res, 39, 580±587. [37] Roome C M, Adam C D (1995), `Crystallite orientation and anisotropic strains in thermally sprayed hydroxyapatite coatings', Biomater, 16, 691±696. [38] Steffens H D, Brune M (1995), `Vacuum-plasma-sprayed titanium±manganese electrode layers for MnO2 deposition', J Therm Spray Techn, 4, 85±88. [39] Disam J, Luebbers K, Neudert K, Sickinger A (1994), `Effect of LPPS spray parameters on the structure of ceramic coating', J Thermal Spray Techn, 3, 142± 147.

ß Woodhead Publishing Limited, 2011

Surface modification by calcium phosphate deposition

161

[40] Kuroda S, Kawakita J, Watanabe M, Katanoda H (2008), `Warm spraying ± a novel coating process based on high-velocity impact of solid particles', Sci Tech Adv Mater, 9, 1±17. [41] Wolke J G C, de Blieck-Hogervorst J M A, Kraak T G, Herlaar W, de Groot K (1991), `The characterization of hydroxyapatite coatings sprayed with VPS, APS and DJ systems', in Proc. 4th Natn. Thermal Spray Conference, Pittsburgh, Pennsylvania, USA, 481±490. [42] Oguchi H, Ishikawa K, Ojima S, Hirayama Y, Seto K, Eguchi G (1992), `Evaluation of a high-velocity flame-spraying technique for hydroxyapatite', Biomater, 13, 471±477. [43] Steinhauser S, Wielage B (1997), `Composite coatings: manufacture, properties and application', Proc 10th Conf Surface Modif Tech, Singapore, 436±450. [44] Khor K A, Li H, Cheang P, Boe S Y (2003), `In vitro behavior of HVOF sprayed calcium phosphate splats and coatings', Biomater, 24, 723±735. [45] Khor K A, Li H, Cheang P (2003), `Processing±microstructure±property relations in HVOF sprayed calcium phosphate based bioceramic coatings', Biomater, 24, 2233±2243. [46] Khor K A, Li H, Kumar R, Cheang P (2004), `Characterization of hydroxyapatite/ nano-zirconia composite coatings deposited by high velocity oxy-fuel (HVOF)', Surf Coat Tech, 182, 227±236. [47] Dhert W J A, Klein C P A T, Wolke J G C, Vandervelde, E A, de Groot K, Rozing P M (1991), `A mechanical investigation of fluorapatite, magnesium whitlockite, and hydroxylapatite plasma-sprayed coatings in goats, J Biomed Mater Res, 25, 1183±1200. [48] Jansen J A, van der Waerden J P C M, Wolke J G C (1993), `Histological investigation of the biological behavior of different hydroxyapatite plasma-sprayed coatings in rabbits', J Biomed Mater Res, 27, 603±610. [49] Sun L M, Berndt C C, Gross K A, Kucuk A (2001),`Material fundamentals and clinical performance of plasma-sprayed hydroxyapatite coatings: a review', J Biomed Mater Res, 58, 570±592. [50] Gross K A, Walsh W, Swarts E (2004), `Analysis of retrieved hydroxyapatitecoated hip prostheses', J Therm Spray Tech, 13, 190±199. [51] Gross K A, Ray N, Rokkum M (2002), `The contribution of coating microstructure to degradation and particle release in hydroxyapatite coated prostheses', J Biomed Mater Res, 63, 106±114. [52] Yang Y Z, Kim K H, Agrawal C M, Ong J L (2003), `Effect of post-deposition heating temperature and the presence of water vapor during heat treatment on crystallinity of calcium phosphate coatings', Biomater, 24, 5131±5137. [53] Yang Y, Bumgardner J D, Cavin R, Carnes D L, Ong J L (2003), `Osteoblast precursor cell attachment on heat-treated calcium phosphate coatings', J Dent Res, 82, 449±453. [54] Rabiei A, Blalock T, Thomas B, Cuomo J, Yang Y, Ong J (2007), `Microstructure, mechanical properties, and biological response to functionally graded HA coatings', Mat Sci Eng C Biol Sci, 27, 529±533. [55] Bauerle D (2000), Laser processing and chemistry, 3rd edn, Berlin, Springer. [56] Willmott P R, Huber J R (2000), `Pulsed laser vaporization and deposition', Rev Mod Phys, 72, 315±328. [57] Singh R K, Narayan J (1990), `Pulsed-laser evaporation technique for deposition of thin-films ± physics and theoretical-model', Phys Rev B, 41B, 8843±8859. [58] Nelea V, Ristoscu C, Chiritescu C, Ghica C, Mihailescu I N, Pelletier H, Mille P,

ß Woodhead Publishing Limited, 2011

162

[59]

[60]

[61] [62] [63] [64] [65] [66] [67] [68] [69]

[70] [71] [72] [73] [74] [75]

Surface modification of biomaterials Cornet A (2000), `Pulsed laser deposition of hydroxyapatite thin films on Ti-5A12.5Fe substrates with and without buffer layers', Appl Surf Sci, 168, 127±131. Nelea V, Pelletier H, Illiescu M, Werckmann J, Craciun V, Mihailescu I N, Ristoscu C, Ghica C (2002), `Calcium phosphate thin film processing by pulsed laser deposition and in situ assisted ultraviolet pulsed laser deposition', J Mater Sci ± Mater Med, 13, 1167±1173. Zeng H T, Lacefield W R, Mirov S (2000), `Structural and morphological study of pulsed laser deposited calcium phosphate bioceramic coatings: influence of deposition conditions, laser parameters, and target properties', J Biomed Mater Res, 50, 248±258. Singh R K, Qian F, Nagabushnam V, Damodaran R, Moudgil B M (1994), `Excimer laser deposition of hydroxylapatite thin films', Biomater, 15, 522±528. Ball M D, Downes S, Scotchford C A, Antonov E N, Bagratashvili V N, Popov V K, Lo W J, Grant D M, Howdle S M (2001), `Osteoblast growth on titanium foils coated with hydroxyapatite by pulsed laser ablation', Biomater, 22, 337±347. Socol G, Torricelli P, Bracci B, Iliescu M, Miroiu F, Bigi A, Werckmann J Mihailescu I N (2004), `Biocompatible nanocrystalline octacalcium phosphate thin films obtained by pulsed laser deposition', Biomater, 25, 2539±2545. Bigi A, Bracci B, Cuisinier F, Elkaim R, Fini M, Mayer I, Mihailescu I N, Socol G, Sturba L, Torricelli, P (2005), `Human osteoblast response to pulsed laser deposited calcium phosphate coatings', Biomater, 26, 2381±2389. Cotell C M, Chrisey D B, Grabowski K S, Sprague J A, Gosset C R (1992), `Pulsed laser deposition of hydroxylapatite thin films on Ti±6Al±4V', J Appl Biomed, 3, 87±93. Cotell C M (1993), 'Pulsed laser deposition and processing of biocompatible hydroxylapatite thin films', Appl Surf Sci, 1993, 227±248. Torrisi L, Setola R (1993), `Thermally assisted hydroxyapatite obtained by pulsedlaser deposition on titanium substrates', Thin Solid Films, 227, 32±36. Craciun V, Boyd I W, Craciun D, Andreazza P, Perriere J (1999), `Vacuum ultraviolet annealing of hydroxyapatite films grown by pulsed laser deposition', J Appl Phy, 85, 8410±8414. Arias J L, Garcia-Sanz F J, Mayor M B, Chiussi S, Pou J, Leon B, Perez-Amor M (1998), `Physicochemical properties of calcium phosphate coatings produced by pulsed laser deposition at different water vapour pressures', Biomater, 19, 883± 888. Mayor B, Arias J, Chiussi S, Garcia F, Pou J, Leon Fong B, Perez-Amor M (1998), `Calcium phosphate coatings grown at different substrate temperatures by pulsed ArF-laser deposition', Thin Solid Films, 317, 363±366. Wang C K, Chern Lin J H, Ju C P, Ong H C, Chang R P H (1997), `Structural characterization of pulsed laser-deposited hydroxyapatite film on titanium substrate', Biomater, 18, 1331±1338. Singh R K, Qian F, Nagabushnam V, Damodaran R, Moudgil B M (1994), `Excimer-laser deposition of hydroxyapatite thin-films', Biomater, 15, 522±528. Hontsu S, Matsumoto T, Ishii J, Nakamori M, Tabata H, Kawai T (1997), `Electrical properties of hydroxyapatite thin films grown by pulsed laser deposition', Thin Solid Films, 295, 214±217. Fernandez-Pradas J M, Sardin G, Cleries L, Serra P, Ferrater C, Morenza J L (1998), `Deposition of hydroxyapatite thin films by excimer laser ablation', Thin Solid Films, 317, 393±396. Nelea V, Morosanu C, Iliescu M, Mihailescu I N (2004), `Hydroxyapatite thin

ß Woodhead Publishing Limited, 2011

Surface modification by calcium phosphate deposition

[76] [77] [78] [79] [80] [81] [82] [83] [84] [85] [86] [87] [88] [89] [90] [91] [92] [93]

163

films grown by pulsed laser deposition and radio-frequency magnetron sputtering: comparative study', Appl Surf Sci, 228, 346±356. Song W-H, Jun Y-K, Han Y, Hong S-H (2004), `Biomimetic apatite coatings on micro-arc oxidized titania', Biomater, 25, 3341±3349. Du X W, Fu Y, Sun J, Yao P (2007), `The effect of annealing atmosphere on photoluminescent properties of SiCN films', Surf Coat Tech, 201, 5404±5407. Wei D, Zhou Y, Wang Y, Jia D (2007), `Characteristic of microarc oxidized coatings on titanium alloy formed in electrolytes containing chelate complex and nano-HA', Appl Surf Sci, 253, 5045±5050. Huang Y, Wang Y, Ning C, Nan K, Han Y (2008), `Preparation and properties of a cerium-containing hydroxyapatite coating on commercially pure titanium by micro-arc oxidation', Rare Metals, 27, 257±260. Han Y, Sun J F, Huang X (2008), `Formation mechanism of HA-based coatings by micro-arc oxidation', Electrochem Comm, 10, 510±513. Cheng H M, Ning C Y, Zheng H D, Tan G X, Ma Q, Zhang D G (2009), `Effect of cathode voltage on structure of micro-arc oxidation films of pure titanium', Rare Metals Mater Eng, 38, 1116±1118. Nan K H, Wu T, Chen J H, Jiang S, Huang Y, Pei G X (2009), `Strontium doped hydroxyapatite film formed by micro-arc oxidation', Mater Sci Eng C-Bio S, 29, 1554±1558. Wolke J G C, van Dijk K, Schaeken H G, de Groot K, Jansen J A (1994), `Study of the surface characteristics of magnetron-sputter calcium phosphate coatings', J Biomed Mater Res, 28, 1477±1484. Wolke J G C, de Groot K, Jansen J A (1998), `Dissolution and adhesion behaviour of radio-frequency magnetron-sputtered Ca-P coatings', J Mater Sci, 33, 3371± 3376. Wolke J G C, de Groot K, Jansen J A (1998), `In vivo dissolution behavior of various RF magnetron sputtered Ca-P coatings', J Biomed Mater Res, 39, 524±530. Wolke J G C, van der Waerden J P C M, Schaeken H G, Jansen J A (2003), In vivo dissolution behavior of various RF magnetron-sputtered Ca-P coatings on roughened titanium implants', Biomater, 24, 2623±2629. Nelea V, Morosanu C, Iliescu M, Mihailescu I N (2003), `Microstructure and mechanical properties of hydroxyapatite thin films grown by RF magnetron sputtering', Surf Coat Tech, 173, 315±322. Chen M, Liu L, You, C, Yang X, Cui, Z (2007), `Interfacial characteristic of graded hydroxyapatite and titanium thin film by magnetron sputtering', Surf Coat Tech, 201, 5688±5691. Xu S, Long J D, Ostrikov K N, Lu J H, Diong C H (2002), `RF magnetron sputtering deposition of bioactive Ca-P-based coatings on Ti-6Al-4V alloy', IEEE T Plasma Sci, 30, 118±119. Nieh T G, Jankowski A F, Koike J (2001), `Processing and characterization of hydroxyapatite coatings on titanium produced by magnetron sputtering', J Mater Res, 16, 3238±3245. Yamashita K, Yagi T, Umegaki T (1996), `Bonelike coatings onto ceramics by reactive magnetron sputtering', J Am Ceram Soc, 79, 3313±3316. Nakamura S, Hamagami J-I, Yamashita K (2007), `Hydrothermal crystallization of carbonate-containing hydroxyapatite coatings prepared by radiofrequencymagnetron sputtering method', J Biomed Mater Res, 80B, 102±106. Feddes B, Wolke J G C, Jansen J A, Vredenberg A M (2003), `Initial deposition of calcium phosphate ceramic on polystyrene and polytetrafluoroethylene by rf

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magnetron sputtering deposition', J Vac Sci Technol A, 21, 363±368. [94] Feddes B, Wolke J G C, Vredenberg A M, Jansen J A (2004), `Adhesion of calcium phosphate ceramic on polyethylene (PE) and polytetrafluoroethylene (PTFE)', Surf Coat Tech, 184, 247±254. [95] Feddes B, Vredenberg A M, Wolke J G C, Jansen J A (2004), `Bulk composition of r.f. magnetron sputter deposited calcium phosphate coatings on different substrates (polyethylene, polytetrafluoroethylene, silicon)', Surf Coat Tech, 185, 346±355. [96] Feddes B, Wolke J G C, Weinhold W P, Vredenberg A M, Jansen J A (2004), `Adhesion of calcium phosphate coatings on polyethylene (PE), polystyrene (PS), poly(tetrafluoroethylene) (PTFE), poly(dimethylsiloxane) (PDMS) and poly-Llactic acid (PLLA)', J Adhes Sci Technol, 18, 655±672. [97] Feddes B, Wolke J G C, Vredenberg A M, Jansen J A (2004), `Initial deposition of calcium phosphate ceramic on polyethylene and polydimethylsiloxane by rf magnetron sputtering deposition: the interface chemistry', Biomater, 25, 633±639. [98] Wan T, Aoki H, Hikawa J, Lee J H (2007), `RF-magnetron sputtering technique for producing hydroxyapatite coating film on various substrates', Biomed Mater Eng, 17, 291±297. [99] Ozeki K, Janurudin J M, Aoki H, Fukui Y (2007), `Photocatalytic hydroxyapatite/ titanium dioxide multilayer thin film deposited onto glass using an rf magnetron sputtering technique', Appl Surf Sci, 253, 3397±3401. [100] Porter A E, Rea S M, Galtrey M, Best S M, Barber Z H (2004), `Production of thin film silicon-doped hydroxyaptite via sputter deposition', J Mater Sci, 39, 1895±1898. [101] Thian E S, Huang J, Best S M, Barber Z H, Bonfield W (2005), `The response of osteoblasts to nanocrystalline silicon-substituted hydroxyapatite thin films', Biomater, 26, 2947±2956. [102] Long J D, Xu S, Cai J W, Jiang N, Lu J H, Ostrikov K N, Diong C H (2002), `Structure, bonding state and in-vitro study of Ca-P-Ti film deposited on Ti6A14V by RF magnetron sputtering', Mater Sci Eng C, 20, 175±180. [103] Xu S Y, Long J D, Sim L N, Diong C H, Ostrikov K (2005), `RF plasma sputtering deposition of hydroxyapatite bioceramics: Synthesis, performance, and biocompatibility', Plasma Process Polym, 2, 373±390. [104] Toque J A, Hamdi M, Ide-Ektessabi A, Sopyan I (2009), `Effect of the processing parameters on the integrity of calcium phosphate coatings produced by rfmagnetron sputtering', Int J Mod Phys B, 23, 5811±5818. [105] Yonggang Y, Wolke J G C, Yubao L, Jansen JA (2007), `The influence of discharge power and heat treatment on calcium phosphate coatings prepared by RF magnetron sputtering deposition', J Mater Sci ± Mater M, 18, 1061±1069. [106] Watazu A, Kimoto K, Tsutomu S, Tanaka K, Sawada T, Toyoda M, Saito N (2007), `Ti-Ca-P films formed by RF magnetron sputtering method using dual targets', Mater Sci Forum, 544±545, 495±498. [107] Boyd A R, Duffy H, McCann R, Cairns M L, Meenan B J (2007), `The Influence of argon gas pressure on co-sputtered calcium phosphate thin film', Nucl Instrum Meth B, 258, 421±428. [108] Boyd A R, Duffy H, McCann R, Meenan B J (2008), `Sputter deposition of calcium phosphate/titanium dioxide hybrid thin films', Mat Sci Eng C ± Bio S, 28, 228±236. [109] Shi J Z, Chen C Z, Yu H J, Zhang S J (2008), `The effect of process conditions on the properties of bioactive films prepared by magnetron sputtering', Vacuum, 83, 249±256. [110] Shi J Z, Chen C Z, Yu H J, Zhang S J (2008), `Application of magnetron sputtering

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[111] [112]

[113] [114] [115] [116] [117]

[118] [119] [120] [121]

[122] [123] [124] [125] [126] [127]

165

for producing bioactive ceramic coatings on implant materials', B Mater Sci, 31, 877±884. Ide-Ektessabi A, Yamaguchi T, Tanaka Y (2006), `Fabrication of hydroxyapatite thin films for biomedical applications using RF magnetron sputtering', Nucl Instrum Meth B, 249, 723±725. Narushima K, Ueda K, Goto T, Katsube T, Kawamura H, Ouchi C, Iguchi Y (2006), `Characterization of calcium phosphate films prepared by RF magnetron sputtering', Materials and Devices for Smart Systems II Symposium, Materials Research Society Symposium Proceedings, Eds Furuya Y, Su J, Takeuchi I, Varadan V K, Ulicny J, 888, 101±106. Narushima T, Ueda K, Goto T, Masumoto H, Katsube T, Kawamura H, Ouchi C, Iguchi Y (2005), `Preparation of calcium phosphate films by radiofrequency magnetron sputtering', Mater Trans, 46, 2246±2252. Ozeki K, Yuhta T, Fukui Y, Aoki H (2002), `Phase composition of sputtered films from a hydroxyapatite target', Surf Coat Tech, 160, 54±61. Ozeki K, Fukui Y, Aoki H (2007), `Influence of the calcium phosphate content of the target on the phase composition and deposition rate of sputtered films', Appl Surf Sci, 253, 5040±5044. O'Kane C, Duffy H, Meenan B J, Boyd A R (2008), `The influence of target stoichiometry on the surface properties of sputter deposited calcium phosphate thin films', Surf & Coat Tech, 203, 121±128. Snyders R, Bousser E, Music D, Jensen J, Hocquet S, Schneider J M (2008), `Influence of the chemical composition on the phase constitution and the elastic properties of RF-sputtered hydroxyapatite coatings', Plasma Process Polym, 5, 168±174. Wang C, Ma J, Cheng W, Zhang R (2002), `Thick hydroxyapatite coatings by electrophoretic deposition', Mater Lett, 57, 99±105. Gao L, Lin J R (2008), `Electrophoretic coating of hydroxyapatite on pyrolytic carbon using glycol as dispersion medium', J Wuhan Univ Technol Mater Sci Ed, 23, 293±297. Monkawa A, Ikoma T, Yunoki S, Ohta K, Tanaka J (2006), `Electrophoretic deposition of hydroxyapatite nanocrystal, Key Eng Mater, 309±311, 643±646. Albayrak O, El-Atwani O, Altintas S (2008), `Hydroxyapatite coating on titanium substrate by electrophoretic deposition method: Effects of titanium dioxide inner layer on adhesion strength and hydroxyapatite decomposition', Surf Coat Tech, 202, 2482±2487. Nie X, Leyland A, Matthews (2000), `A Deposition of layered bioceramic hydroxyapatite/TiO2 coatings on titanium alloys using a hybrid technique of micro-arc oxidation and electrophoresis', Surf Coat Tech, 125, 407±414. Xiao X F, Liu R F, Tang X L (2009), `Electrophoretic deposition of siliconsubstituted hydroxyapatite/poly(epsilon-caprolactone) composite coatings', J Mater Sci ± Mater M, 20, 691±697. Peaker A J S, Czernuszka J T (1996), `The effect of electric field on the formation of hydroxyapatite coatings', Thin Solid Films, 287, 174±183. Javidi M, Javadpour S, Bahrololoom M E, Ma J (2008), `Electrophoretic deposition of natural hydroxyapatite on medical grade 316L stainless steel', Mat Sci Eng C ± Bio S, 28, 1509±1515. Pang X, Zhitomirsky I (2007), `Electrophoretic deposition of composite hydroxyapatite-chitosan coatings', Mater Character, 58, 339±348. Montero-Ocampo C, Villegas D, Veleva L (2008), `Electrochemical assisted

ß Woodhead Publishing Limited, 2011

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[128] [129] [130] [131] [132] [133] [134] [135] [136] [137] [138] [139] [140] [141] [142] [143] [144]

Surface modification of biomaterials deposition of calcium phosphate coatings for orthopaedic applications', Surf Eng, 24, 18±22. Ma M, Ye W, Wang X X (2008), `Effect of supersaturation on the morphology of hydroxyapatite crystals deposited by electrochemical deposition on titanium', Mat Lett, 62, 3875±3877. Hu H B, Lin C J, Lui P P Y, Leng Y (2003), `Electrochemical deposition of hydroxyapatite with vinyl acetate on titanium implants', J Biomed Mater Res, 65A, 24±29. Wang Y Q, Tao J, Wang L, He P T, Wang T (2008), `HA coating on titanium with nanotubular anodized TiO2 intermediate layer via electrochemical deposition', T Nonferr Metal Soc, 18, 631±635. Blackwood D J, Seah K H W (2009), `Electrochemical cathodic deposition of hydroxyapatite: improvements in adhesion and crystallinity', Mat Sci Eng C ± Bio S, 29, 1233±1238. Roessler S, Sewing A, Stoelzel M, Born R, Scharnweber D, Dard M, Worch H (2003), `Electrochemically assisted deposition of thin calcium phosphate coatings at near-physiological pH and temperature', J Biomed Mater Res, 64A, 655±663. Wang J W, de Boer J, de Groot K (2009), `Proliferation and differentiation of osteoblast-like MC3T3-E1 cells on biomimetically and electrolytically deposited calcium phosphate coatings', J Biomed Mater Res, 90A, 664±670. Hu H B, Lin C J, Hu R, Leng Y (2002), `A study on hybrid bioceramic coatings of HA/poly(vinyl acetate) co-deposited electrochemically on Ti-6Al-4V alloy surface', Mat Sci Eng C ± Bio S, 20 SI, 209±214. Li M Q, Qu L J, Ma C, Yang S Q (2008), `Bioactive and stability of calcium phosphate-polypyrrole composite coatings by electrochemical deposition', Key Eng Mater, 368±372, 1198±1200. Redepenning J, Venkataraman G, Chen J, Stafford N (2003), `Electrochemical preparation of chitosan/hydroxyapatite composite coatings on titanium substrates', J Biomed Mater Res, 66A, 411±416. Sun T, Wong J H W, Wang M (2009), `Fabrication of an apatite/collagen composite coating on the NiTi shape memory alloy through electrochemical deposition and coating characterisation', Mater Sci Forum, 618±619, 319±323. Lopatin C M, Pizziconi V, Alford T L, Laursen T (1998), `Hydroxyapatite powders and thin films prepared by a sol-gel technique', Thin Solid Films, 326, 227±232. Wang D G, Chen C Z, Ting H, Lei T Q (2008), `Hydroxyapatite coating on Ti6Al4V alloy by a sol-gel method', J Mater Sci ± Mater M, 19, 2281±2286. Piveteau L D, Girona M I, Schlapbach L, Barboux P, Boilot J, Gasser B (1999), `Thin films of calcium phosphate and titanium dioxide by a sol-gel route: a new method for coating medical implants', J Mater Sci ± Mater M, 10, 161±167. Azem F A, Cakir A (2009), `Synthesis of HAP coating on galvanostatically treated stainless steel substrates by sol-gel method', J Sol-Gel Sci Techn, 51, 190±197. Wie H, Hero H, Solheim T, Kleven E, Rorvik A M, Haanaes H R (1995), `Bonding capacity in bone of HIP-processed Ha-coated titanium-mechanical and histological investigations', J Biomed Mater Res, 29, 1443±1449. Khor K A, Yip C S, Cheang P (1997), `Post-spray hot isostatic pressing of plasma sprayed Ti-6Al-4V/ hydroxyapatite composite coatings', J Mater Process Tech, 71, 280±287. Onoki T, Hashida T (2006), `New method for hydroxyapatite coating of titanium by the hydrothermal hot isostatic pressing technique', Surf Coat Tech, 200, 6801± 6807.

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[145] Onoki T and Hashida T (2006), `New method for hydroxyapatite coating by hydrothermal hot-pressing via double layered capsule', Key Eng Mater, 309±311, 647±650. [146] Tanahashi M, Yao T, Kokubo T, Minoda M, Miyamoto T, Nakamura T, Yamamuro T (1994), `Apatite coating on organic polymers by a biomimetic process', J Am Ceram Soc, 77, 2805±2808. [147] Abe Y, Kokubo T, Yamamuro T (1990), `Apatite coating on ceramics, metals and polymers utilizing a biological process', J Mater Sci ± Mater M, 1, 233±238. [148] Miyaji F, Kim H M, Handa S, Kokubo T, Nakamura T (1999), `Bonelike apatite coating on organic polymers. Novel nucleation process using sodium silicate solution', Biomater, 20, 913±919. [149] Taguchi T, Muraoka Y, Matsuyama H, Kishida A, Akashi M (2001), `Apatite coating on hydrophilic polymer-grafted poly (ethylene) films using an alternate soaking process', Biomater, 22, 53±58. [150] Oliveira A L, Malafaya P B, Reis R L (2003), `Sodium silicate gel as a precursor for the in vitro nucleation and growth of a bone-like apatite coating in compact and porous polymeric structures', Biomater, 24, 2575±2584. [151] Kim HM, Uenoyama M, Kokubo T, Minoda M, Miyamoto T, Nakamura T (2001), `Biomimetic apatite formation on polyethylene photografted with vinyltrimethoxysilane and hydrolyzed', Biomater, 22, 2489±2494. [152] Oyane A, Kawashita M, Kokubo T, Minoda M, Miyamoto T, Nakamura T (2002), `Bonelike apatite formation on ethylene-vinyl alcohol copolymer modified with a silane coupling agent and titania solution', J Ceram Soc Jpn, 110, 248±254. [153] Oyane A, Kawashita M, Nakanishi K, Kokubo T, Minoda M, Miyamoto T, Nakamura T (2003), `Bonelike apatite formation on ethylenevinyl alcohol copolymer modified with silane coupling agent and calcium silicate solutions', Biomater, 24, 1729±1735. [154] Kokubo T, Kushitani H, Sakka S, Kitsugi T, Yamamuro T (1990), `Solutions able to reproduce in vivo surface-structure changes in bioactive glass±ceramic A-W', J Biomed Mater Res, 24, 721±734. [155] Kokubo T, Kim H M, Kawashita M, Nakamura T (2001), `Process of calcification on artificial materials', Z Kardiol, 90, 86±91. [156] Juhasz J A, Best S M, Bonfield W (2010), `Preparation of novel bioactive nanocalcium phosphate±hydrogel composites', Sci Technol Adv Mat, 11, 014103. [157] Oliveira A L, Elvira C, Vasquez B, San Roman J, Reis R L (1999), `Surface modifications tailor the characteristics of biomimetic coatings nucleated on starch based polymers', J Mater Sci ± Mater Med, 10, 827 [158] Oliveira A L, Gomes M E, Malafaya P B, Reis R L (2003), `Biomimetic coating of starch based polymeric foams produced by a calcium silicate based methodology', Key Eng Mater, 240±242, 101±104. [159] Kim S S, Park M S, Gwak S J, Choi C Y, Kim B S (2006), `Accelerated bonelike apatite growth on porous polymer/ceramic composite scaffolds in vitro', Tissue Eng, 12, 2997±3006. [160] Liu Y, Hunziker E B, Randall N X, de Groot K, Layrolle P (2003), `Proteins incorporated into biomimetically prepared calcium phosphate coatings modulate their mechanical strength and dissolution rate', Biomater, 24, 65±70. [161] Wittmer C R, Phelps J A, Saltzman W M, van Tassel PR (2007), `Fibronectin terminated multilayer films: protein adsorption and cell attachment studies', Biomater, 28, 851±860. [162] Chen W, Liu Y, Courtney H S, Bettenga M, Agrawal C M, Bumgardner J D, Ong J

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[163] [164] [165] [166] [167] [168] [169] [170]

Surface modification of biomaterials L (2006), `In vitro anti-bacterial and biological properties of magnetron cosputtered silver-containing hydroxyapatite coating', Biomater, 27, 5512±5517. Brohede U, Forsgren J, Roos S, Mihranyan A, Engqvist H, Stromme M (2009), `Multifunctional implant coatings providing possibilities for fast antibiotics loading with subsequent slow release', J Mater Sci ± Mater M, 20, 1859±1867. Zawadzak E, Bil M, Ryszkowska J, Nazhat S N, Cho J, Bretcanu O, Roether J A, Boccaccini A R (2009), `Polyurethane foams electrophoretically coated with carbon nanotubes for tissue engineering scaffolds', Biomed Mater, 4, 15±24. Juhasz J A, Best S M, Auffret A D, Bonfield W (2008), `Biological control of apatite growth in simulated body fluid and human blood serum', J Mater Sci ± Mater M, 19, 1823±1829. Zhao, J, Cai, X M, Tang, H Q, Liu T, Gu H Q, Cui R Z (2009), `Bactericidal and biocompatible properties of TiN/Ag multilayered films by ion beam assisted deposition', J Mater Sci ± Mater M, 20 SI, 101±105. Ye G and Troczynski T (2008), `Hydroxyapatite coatings by pulsed ultrasonic spray pyrolysis', Ceram Int, 34, 511±516. Leeuwenburgh, S C G, Wolke J G C, Siebers T C, Schoonman J, Jansen, J A (2006), `In vitro and in vivo reactivity of porous, electrosprayed calcium phosphate coatings', Biomater, 27, 3368±3378. Aguilar-Frutis, M, Kumar, S, Falcony, C (2009), `Spray-pyrolyzed hydroxyapatite thin-film coatings', Surf Coat Tech, 204, 1116±1120. Marti, P C (2009), `Zinc-containing hydroxyapatite coatings for orthopaedic applications', PhD Thesis, University of Cambridge.

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Plate II Magnetron-sputtered HA coating on titanium alloy substrate (a) showing the grooves due to grinding of the underlying substrate and fluorescence microscopy showing the actin cytoskeleton (red) and nuclear DNA (blue) of HOBs grown on HA-coated substrate for 3 days (b). From Marti PC (2009) `Zinc-containing hydroxyapatite coatings for orthopaedic applications', PhD Thesis, University of Cambridge.

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Biomaterial surface topography to control cellular response: technologies, cell behaviour and biomedical applications V . R . K E A R N S , University of Liverpool, UK and R . J . M c M U R R A Y and M . J . D A L B Y , University of Glasgow, UK

Abstract: The chapter discusses the use of substrate topography to control cellular response. It details some of the fabrication methods used to produce controlled micro- and nanotopography, including both bottom-up and topdown techniques. The mechanisms by which cells respond to surface topography are discussed, in particular highlighting the complexity of this response and the importance of scale. The use of topography to control stem cell differentiation is also considered, and some of the potential biomedical and tissue engineering applications of surface topography are identified. Key words: surface topography, nanotopography, stem cell differentiation, cell orientation.

7.1

Introduction

In their natural environment, cells will experience complex chemical and topographical cues. Cells may encounter different sizes of topography, from macro, such as bone or ligament shape, to micro, such as the shapes of other cells, to nano, such as basement membranes, protein folding and collagen banding (64 nm repeat pattern). Cells themselves exhibit nanoscale features, such as filopodia. Interactions between the cells and chemically and topographically patterned extracellular matrix (ECM) via cell adhesion molecules are crucial in determining cell proliferation, migration, differentiation and death. Thus, understanding the effect of surface topography on cellular response is crucial, particularly for tissue engineering. In vitro, cells will be exposed to a different environment but this still has the capacity to affect cell behaviour. It has, in fact, been known for almost 100 years that cells will respond to the shape of their environment [1]. The idea of studying specifically the response of the cultured cell to shape was then picked up in the 1960s by Curtis [2]. In the 1980s photolithographical techniques from the microelectronic industry were developed to allow the patterning of areas large

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enough for cell culture and analysis and this led to a raft of research that demonstrated that every cell type tested would respond to micron-scale topography by adjusting adhesion, migration, morphology, cytoskeleton and, more recently, the genome. In the meantime, driven by the need to fabricate faster microchips, engineers were developing lithographical techniques to allow the placement of increasing numbers of transistors on chips. This lead to the development of electron beam lithography and its development for use with cultured cells [3]. These top-down fabrication techniques have helped develop new understanding of the limits of cell response to the shape of their environment. Cell structures, such as filopodia, have been shown, in fact, to guide contact to features as small as 10 nm in height [4]. In this chapter, we will describe methods of manufacturing controlled surface topography, discuss the various mechanisms by which cells respond to these surfaces and identify some potential applications of this ability to control the cellular response.

7.2

Defining micron and nano

It is important to define the scale of features to be discussed. In this chapter, we will define surface features that have one dimension greater than 1 m as `micron-scale' and features less than 100 nm as `nanoscale'. Those features that are less than 1 m but greater than 100 nm will be described as `sub-micron'. However, we would like to caution the reader that many authors refer to features less than 1 m as `nanoscale', and suggest that particular attention must be paid to the terminology used by different investigators.

7.3

Manufacturing surface topography

Nanotopography is only very recently becoming available to the biologist for study due to the prohibitive expense of producing large areas of defined nanofeatures. However, recent advances in electronic engineering, physics and chemistry are making substrates with controlled dimensions more readily available. We start with photolithography (PL) as it is necessary to understand the principles of lithography employed by engineers in order to see the key developments towards producing cheaper nanofeatures over larger areas.

7.3.1

Photolithography (PL)

The microchip industry has traditionally used PL to fabricate micronscale designs. In order to do this, a light sensitive substrate (photoresist) is spread as a thin film onto silicon wafers (Fig. 7.1(a)). Light can then be directed onto the resist through a patterned mask, exposing the photosensitive layer in a precise

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7.1 Two types of lithography. (a) Photolithography ± a mask is used to expose a radiation-sensitive resist. This can then be developed and reactive ion etched to form a topography. Electron beam lithography is a development of this technique where electrons are used rather than light. (b) Colloidal lithography ± nanocolloids are dispersed across the surface of a material and then used as an etch mask for reactive ion etching. Figure used with permission from Elsevier ± Dalby et al., Use of nanotopography to study mechanotransduction in fibroblasts ± methods and perspectives. Eur J Cell Biol 83, 159±69 (2004).

pattern. The resist is then developed (analogous with photography), and the nonexposed or exposed area (depending on the resist used) can be removed, thus leaving a topography of resist around areas of exposed silicon. The resist is then baked to make it hard and the substrate is ready for etching. Etching can be wet (chemical) or dry (reactive ion bombardment), both having the effect of simultaneously removing the hard resist and etching into the substrate. This leaves the samples resist free and with a topographical pattern. In

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electronics, doping can take place to give electrical properties to the substrate. For biological purposes, it is desirable to stop after etching and grow cells on the patterned material. It is also possible at this stage to emboss, or cast, the pattern into a polymer suitable for cell growth. Whilst PL was used by many of the pioneers of cell response to microtopography [5±7], its usefulness for producing nanotopography is limited. Thus, alternative fabrication techniques have been sought.

7.3.2

Laser holography

Clark et al. [8] developed an interference technique used to produce ultrafine masks for X-ray printing. Samples coated in photoresist can be exposed to patterns produced in a two-mirror interferometer, producing a linear light pattern whose period is given by: period ˆ =2 sin  ( ˆ light wavelength of light,  ˆ half angle of the recombination of the beams). Development of the photoresist (as with PL) results in the formation of an etch mask. Dry etching can then be used to give depth to the pattern. Patterns produced by laser holography with a 260 nm groove width and 100 nm depth have been shown to induce BHK fibroblast, MDCK epithelial, and primary neurite contact guidance. This technique, however, is still limited in terms of the variety of pattern that can be produced, the smallest size of the dimensions that can be made and cost.

7.3.3

Electron-beam lithography

In order to achieve even better resolution, electron beams are now being used as a radiation source for photoresist development. In electronics, the smaller devices can be made, the more engineers can fit on a chip and the faster they can go; this has led to Moore's law where the density of transistors on a chip has been seen to double every 1.9 years since the 1960s [9]. Thus, there is a massive push at present to develop nanochips. This necessitates the use of precision nanofabrication. At present, electron beam lithography (e-beam) offers the best precision and resolution (down to 510 nm in X and Y axes, less than 1 nm in the Z axis). For this technique, computer aided design software is used to control electron deflection and exposure time onto a radiation sensitive resist (thus rather than shining a `block' of light through a mask exposing all areas to be patterned in parallel, the electrons are guided with great precision onto the resist directly, and the pattern written serially). The resist is then developed and etched to produce a surface topography in a similar way to that described for PL. The use of this technique in biological applications has been limited due to distortion of the beam when deflected by more than 1 mm limiting the size of the patterns produced. Recently, techniques using interferometric control have been developed to allow movement of the sample in the e-beam writer and thus, stitching of the patterns into large areas [3]. This development has allowed for

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the production of large enough areas of pattern to be produced for cell testing; thus far, cell adhesion and filopodial interaction have been measured, with notable decreases in focal contact formation and numbers of cells adhered [10]. E-beam is, however, expensive and time consuming due to serial exposure defining the pattern, and the more complex the pattern, the more complex the programming and the more time taken. It is interesting to note that researchers are now also able to produce biomimetic designs using e-beam, for example reproducing collagen banding [11]. However, fabricators are now becoming successful in producing much cheaper nanotopography, over larger areas. Such techniques (including colloidal lithography) are potentially very useful to biologists due to their low costs. For more details on electron-beam lithography and its applications, the reader is directed to [12].

7.3.4

Colloidal lithography

This technique uses nanocolloids as an etch mask, and can be used to produce large areas of nanotopography with short-range order (Fig. 7.1(b)). The colloids are dispersed as a monolayer and electrostatically self-assembled on a surface (i.e., the inter-colloidal spacing is altered by changing the charge on the particles). The colloids are then etched away, as is the surrounding substrate material by directed reactive ion bombardment [13, 14]. If the etch is directly from above, this technique produces nanocolumns; by altering the angle of etch, nanospheres and nanocones may also be produced.

7.3.5

Polymer demixing

As well as `top-down' fabrication techniques, some `bottom-up' methodologies have also been used to produce surfaces for cell studies. Polymer demixing relies on the spontaneous phase separation of polymer blends and the demixing of island topographies allows the production of single chemistry substrates, ideal for cell testing as the complexities of mixed surface chemistries can be ignored. The surface of a solid amorphous polymer, or polymer mixture, is expected to be smooth and featureless. However, under certain circumstances marked topography can be obtained from a spontaneous demixing of the components of a binary blend, i.e. in thin films of certain polymer mixtures on selected substrates. This method also allows the production of islands with controlled heights from 10 to 100 nm, depending on the polymer blend used. The factors that determine demixing have been extensively studied. Principally, the mutual compatibility of the polymers controls the behaviour. The relationship between polymer compatibility and generation of topography in thin films of the blend is illustrated by examining a blend of poly(styrene) with a series of brominated polymers, poly(p-bromox-styrene), where x is the fraction of aromatic rings that are brominated, 1  x  0. The compatibility of the

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brominated polymer with poly(styrene) depends on the extent of bromination [15]. When x ˆ 1 the polymers are incompatible and the compatibility increases as x ! 0. A measure of the compatibility is the thickness of the interfacial boundary layer between a poly(styrene) film and a poly(p-bromox-styrene) overlayer. Neutron reflectivity data showed a monotonic correlation between the interfacial layer thickness and the extent of bromination of the bromopolymer [16]. Thin films of blends of the bromopolymer and polystyrene have smooth surfaces at low values of x. As x increases topography appears, becoming more marked with increase in x, corresponding to the decreasing compatibility. The type of topography depends on the relative amounts of the polymer components of the blend [15]. In the poly(styrene)±poly(p-bromostyrene) system, islands of the brominated polymer protrude from a sea of the poly(styrene) when the weight fraction of the bromopolymer is less than ca. 60%. At higher weight fractions the bromopolymer islands link up to form a sheet with holes that expose the underlying poly(styrene). Similar changes in topography have been observed with other binary polymer mixtures. The topography can be controlled further by changing the overall thickness of the polymer film. With spin-cast samples, this is readily accomplished by varying the concentration of the polymer blend solution and/or the spin speed. In general, increase in thickness of the cast films results in islands that are higher and wider. A concomitant change in the number density of the features is observed. More subtle control of the topography may be obtained by varying the molecular weights of the polymer components [17]. From the point of view of producing materials for study of cell interactions, it is desirable to reduce the experimental parameters by having the cells exposed to only one chemistry; the poly(styrene)±poly(p-bromostyrene) system produces a surface with patches of poly(styrene) and patches of poly(p-bromostyrene). In this case, the surface chemistry can be changed by annealing the films [16]. Heating the film to the glass transition temperature of the poly(styrene) causes the poly(styrene) to migrate over the surface, forming an overlayer, thus achieving topography with one chemistry at the surface. The topography is altered by this process, which typically causes the islands to grow in height and width. However, a minimum annealing gives a thin overlayer of poly(styrene) with little change in the topography. Another polymer mixture, which gives similar island topography to the above, is poly(styrene)-poly(n-butylmethacrylate) (PnBMA) [17]. In this system an overlayer is formed at room temperature. The cast film develops topography that is covered spontaneously by a thin overlayer of the acrylate. The process of formation of the demixed polymer film is complex, involving the compatibility of the polymer components, their air-polymer and polymersubstrate surface tensions, polymer component molecular weight, solvent used, rate of solvent evaporation, etc. The film is generally also metastable and can change if the temperature is sufficient to allow movement of one, or both,

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polymer components. These factors have been considered in [16±18] and references therein. More recently, developments in phase separation with block co-polymer systems have allowed the fabrication of nanosurfaces with order approaching that achievable by e-beam lithography [19].

7.4

How surface topography affects cell behaviour

7.4.1

Importance of feature size

There is some dispute as to which size of features has the most significant effect on cell behaviour. The magnitude of the effect is often characterised by orientation, which, of course, only applies to anisotropic, groove-like surface features, although other measures, such as attachment [20], proliferation [21] and phagocytosis [22] have also been investigated. The majority of studies report that increasing feature size increases orientation, up to a limit above which cells do not respond to the topography [23±25], although a few studies disagree [26]. Differences are probably due to the different mechanisms by which surface features exert their effect (discussed later in this chapter) and variety in reported feature dimensions (width, depth and pitch), and thus it can be difficult to summarise the effect of feature dimensions on cell behaviour. Recently, a systematic study suggested that the aspect ratio of topographical features is a more comprehensive way of being able to study the effect of features on cell alignment [27]. As part of this study, the work of other investigators was reanalysed and, despite large differences in cell types and feature dimensions, a similar correlation between cell alignment and feature aspect ratio is revealed. Potentially, this is a way of being able to identify trends from a wide body of literature. Generalisation of the effect of cellular response to isotropic topography is also complicated by conflicting reports. Several reports describe increased proliferation on micron-scale and sub-micron features compared to smooth controls [28±30], whereas others report no difference or the opposite trend [31±33]. A synergistic effect between micron- and submicron-scale topography on osteoblast proliferation has also been reported [34]. There may also be an upper and lower limit, even for nanoscale roughness, outside which cell viability is compromised [35]. Despite conflicting evidence, some generalisations can be made. There is rapidly growing evidence that feature size and cell adhesion/response may be tuneable in the nanoscale. Nanoislands produced by polymer demixing are a good exemplar [36]. Very small islands (< 20 nm) promote adhesion in most cell types tested (endothelial [37], fibroblastic [38] and mesenchymal stem [39]). As the islands increase in size, they become less adhesive. Islands produced from a blend of polystyrene/polybutylmethacrylate were adhesive to fibroblasts at 10 nm high, but almost completely non-adhesive at 50 nm high [40]. Topographies produced by colloidal lithography have shown a similar trend, with very small

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nanocolumns (11 nm high) producing increased osteoprogenitor response [39], whilst larger nanocolumns (160 nm high) produced a significant (but far from complete) reduction in fibroblast adhesion [41, 42]. Some interesting work on topographies fabricated through the colloidal lithography route showed 100 nm high columns reduce epithelial cell adhesion [43] and, similar to the polymer demixed nanotopographies, that increasing column size decreased epithelial cell adhesion [44]. This work also showed that the 160 nm high nanocolumns imparted properties of low-adhesion on osteoblasts [45]. Studies using another fabrication method, that of interference lithography, also agree with these observations of size and adhesion. Sharp nanoposts (pointed columns) and nanogratings (very sharp grooves) of different heights in 3 size ranges (50±100 nm high, 200±300 nm high and 500±600 nm high) have been demonstrated to decrease adhesion and spreading as size increased [46]. The notable observation from the above is the theme of tuneable height of features, with not much mention about depth. However, there is a very low adhesion pit system that with some cell types can produce almost total nonadhesion. That of nanopits fabricated in a square array by e-beam lithography with diameters of approximately 100±150 nm and centre-centre spacing (pitch) in the range of 300 nm. Again, these adhesion effects have been shown to be size-dependent. A study that used 3 pit sizes (diameter (nm) : pitch (nm) ± 35 : 100, 75 : 200 and 120 : 300) and fibroblasts as a cell model showed that as pit diameter increased, filopodial interaction increased, but cell spreading decreased, i.e. the interaction of the filopodia with the pits was preventing adhesion [47]. A previous study with pit diameters of 150 nm showed almost no epitenon attachment [10]. However we know that large pits (in the range produced by photolithography) promote increased cellular response in, e.g., mesenchymal stem cells [48]. Thus, there is a large gap in our knowledge of the transition between non-adhesive and adhesive pits. This is further complicated by pit symmetry being implicated as important. This was originally mooted by Curtis and others in 2004 [49]. Subsequent studies have used 120 nm diameter, 300 nm pitch pits in both square and hexagonal arrangement. Both with mesenchymal stem cells [50] and with fibroblasts [47, 51, 52], the hexagonal arrangement has been shown to be the least adhesive [53].

7.4.2

Mechanisms of cellular response

In order to best exploit the effect of surface topography on cellular response, it is necessary to understand how the interactions occur. It is unlikely that there will be a `one size fits all' explanation; indeed there are different competing and complementary theories, depending on the nature and size of the surface features. O'Hara and Buck proposed that grooves affected the location of focal adhesions, which in turn would result in a cell oriented to the topography [54].

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Clark and co-workers [55, 56] disputed this on the grounds that different step heights elicited a different cell response and that steps of different heights were unlikely to result in differences in focal adhesion formation. Additionally, Oakley and Brunette observed focal adhesions oriented perpendicular and parallel to the same surface groove or ridge [57], which would not provide a consistent signal to the cell. Clark et al. were one of the first groups to propose a mechanism for this phenomenon based on the physical barrier to cell spreading presented by grooves, or steps. As cells migrate across a substrate they extend filopodia to explore the substrate and form local adhesions. The theory suggests that limitations in the flexibility of the cell cytoskeleton mean that the probability of a cell being able to make a local adhesion when a step is encountered is low, meaning it is unlikely that the cell will overcome the obstacle presented by steps. This results in a high frequency of cells migrating and aligning in the direction with no obstacle. This model explains the fact that cells on surfaces with wider surface features are less likely to be aligned as the cytoskeletal deformation required is less extreme, and thus cells are more likely to be able to overcome the obstacle presented by the features. However, grooves that are of a similar scale to the size of the cell are ineffective as a barrier [58] or an orienting stimulus [59], as are grooves that are relatively narrow such that the cells can bridge the gaps [60]. One limitation of the `actin cytoskeleton' theory is that microtubules appear to be more sensitive than actin filaments to surface topography. Studies by Oakley and Brunette show that microtubules align to surface grooves earlier than actin microfilaments [57] and can cause cell alignment in the absence of an intact actin cytoskeleton [61]. Furthermore, this mechanism does not explain fully the response of cells to surfaces features that do not provide a so-called `continuous edge', such as surface pits and tubes. A study of the cellular response to discontinuous micron-scale features found that in contrast to the response to grooved surfaces, focal adhesion formation at feature gaps preceded the localisation of microtubules and actin microfilaments to the same features [62]. This supports the suggestion [63] that there are different mechanisms for cellular response to topography. In fact, we postulate that the differences in observations are probably related to scale. At micron-scale heights similar to that of the cell, the cells will be forced to orient to the features (in a top-down fashion, where the guidance cues are big and orientate the small cells). However, at the nanoscale, contact guidance is more likely to occur one adhesion at a time. Cells can align to very small step heights and it takes longer for cells to align as the height decreases. In tissues, individual cells align the extracellular matrix directly underneath them and in a small adjacent area through remodelling of focal adhesions. This alignment a little beyond themselves will help to orientate surrounding cells through adhesion alignment at the nanoscale. This considered, it appears that nanoscale cell guidance could work in a bottom-up manner with individual

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adhesions aligning at the nanoscale and eventually aligning the much larger cell. Compelling evidence for this can be found in [64]. So far, we have discussed groove topographies and focal adhesions. In fact, many other shapes at micro- and nano-scale can influence cells. At the micron level, there are, again, perhaps predictable trends. Large pits will entice and entrap cells [65]. At the nanoscale, we perhaps have to consider cell features that are of the same scale as the topography. The cell is an order of magnitude larger ± mesenchymal stem cells can be up to 400±500 m long, yet they can interact with features as small as 10 nm in size. Cell receptors (e.g. integrins) and cytoskeleton are far more on the length-scale of nanotopogarphy. Cell filopodia are membrane projections driven by microfilaments and containing integrin receptors. They also have a tip diameter of around 100 nm. Importantly, they have been seen to interact with features down to 10 nm high by aligning to the features (contact guidance at the sub-cellular level) [47, 66] (Fig. 7.2). Filopodial `sensing' has been implicated in cell exploration on both adhesive [4] and nonadhesive surfaces [67]. Thus, it is clear that focal contacts have an important role in cell alignment to surface topography, suggesting an influence on the actin cytoskeleton. Changes in the cytoskeleton in response to surface topography occur soon after seeding [22, 57] and in addition to changes in focal adhesion orientation, the size of focal adhesions may be affected by the size of the topographical feature [24]. Focal adhesions have an important role in the regulation of intracellular signal transduction (reviewed in [68]) and thus the control of focal adhesions by substrate topography discussed above in addition to the changes in the cell cytoskeleton may influence a variety of cellular functions. Cell shape is important for a variety of cell functions; modification of cell shape can be used to regulate proliferation or apoptosis [69]. It has been proposed that mechanical influences due to topography can be transduced to the cell cytoskeleton via Rho (a member

7.2 Filopodial interaction with nanosubstrates; note the similar length-scales between the cell projections and the nanofeatures. (a) 160 nm high columns originally fabricated by colloidal lithography, (b) 10 nm high islands originally fabricated by polymer demixing.

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of the Rho GTPase family) and its downstream effector, Rho kinase (ROCK). Proteins in the Rho GTPase family regulate the signalling pathways involved in organisation of the cell cytoskeleton and are involved in the regulation of gene transcription [70]. Rho has been identified as a critical factor in the control differentiation of human mesenchymal stem cells (hMSC) [71]. In one study, inhibition of RhoA resulted in adipogenesis whereas increasing RhoA activity caused osteogenesis [72]. The control of RhoA activity and thus cell shape was achieved via chemical patterning and could override the effect of soluble factors in the media (i.e., cell shape could result in cells of osteogenic lineage when grown in adipogenic medium). In those studies, the cell shape was modified by use of a micron-scale chemical patterning of the surface, but it is easy to see how topographical control of cell shape could result in a similar control of cell behaviour. This can be tied in more closely with topography when considering downstream effectors of RhoA. Activation of G-proteins, such as Rho, Rac and Cdc42 will drive actin contractility, filopodial and lamellipodial extension and are mediated, partly by focal adhesion kinase ± a part of the cells focal adhesion. These changes will also feed into the mitogen activated protein kinase pathways, specifically ERK. ERK has now been postulated a central biochemical target for topographical interventions at the microscale [73] and nanoscale [74±77]. We would suggest that perhaps parts of all the theories discussed above are true. It is easy to see how a large step cue can influence filipodial extension and act as a barrier to adhesion formation at the microlevel. This is rather more difficult to visualise at the nanolevel. Here, it is more likely that, for example, contact guidance will happen one adhesion at a time. Nanogrooves (that the cells can grow over) will orientate cells, but with less efficiency. On the groove ridges the adhesions are seen to align mainly in parallel to the ridges, but some also perpendicular [77]. Many reports of cells in grooves report parallel alignment of actin, but recently there have been observations of perpendicular actin arrangements in well-aligned cells. So far, we have considered indirect, biochemical mechanotransduction. Another form of mechanotransduction, direct, is considered to act on the cytoskeletons as an integrated unit [78]. An interesting theory is that of cellular tensegrity, whereby the cells' mechanical structure is explained via tensional integrity [79±85]. Through this tensegrity structure, tensional forces from the extracellular environment (e.g., from tissue loading or changes in cell spreading) are possibly conferred to the nucleus and alter transcriptomic regulation [78, 86±89]. For topography, changes in cell morphology arise from changes in adhesion and subsequent cytoskeletal tension. Tensegrity or not, the inhomogeneity of the cytoskeleton is probably central to the ability of cells to convey tension to the nucleus [90]. We currently support the idea that the intermediate filaments of the cytoskeleton are linked to the lamin intermediate filaments of the nucleoskeleton. It

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7.3 Changes in nucleus morphology in response to low-adhesion nanotopography. (a) Well-spread fibroblast nuclei with centromeres for chromosome 3 stained by FISH. (b) Fibroblast nuclei with relaxed morphology after culture on hexagonal nanopattern (inset ± 120 nm diameter pits originally fabricated by electron beam lithography, image courtesy of Dr Nikolaj Gadegaard, University of Glasgow, UK), showing changed relative arrangement (i.e. closer together) of the centromeres.

is known that the telomeric ends of the interphase chromosomes are intimately linked to the peripheral lamins [91]. Thus, tension directed through the cytoskeleton may be passed directly to the chromosomes during gene transcription. Changes in chromosomal three-dimensional arrangement (Fig. 7.3) may affect transcriptional events such as access to the genes by transcription factors and polymerases. Also, changes in DNA tension can also cause polymerase enzymes to slow down, speed up or even stall completely [92] and may affect proximity to transcription factories [93]. At this point, it is important to mention another recent theory of how nanotopography may alter cytoskeletal mechanotransductive events, that of nanoimprinting into cells by nanofeatures [94, 95]. This describes a phenomenon that has been clearly seen in platelets and for which some evidence has been provided in more complex cell types. For nanoimprinting to occur, the pattern of the topography must be transferred to the cytoskeletal filaments, i.e the topographies produce a template that is favourable/unfavourable to condensation of cytoskeletal polymer chains through invagination of the basal membrane against the topography. Recent study has shown that 3 integrin may be involved in this process [96]. There is evidence that this leads to increased `attempted' endocytosis, i.e. the cells recognise the features as being in the correct size range to try to endocytose and to form claterin coated pits [97]. Such endocytotic vesicles are moved by actin cables ± perhaps it is this mechanism that is causing the topography-mimicking actin patterning described by Curtis and others (Fig. 7.4).

7.4.3

Protein adsorption

Surface topography may affect cell behaviour, particularly alignment, by altering protein adsorption. It is well known that the surface properties of every

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7.4 Nanoimprinting into fibroblasts. (a) Lamellae of a human fibroblast with 160 nm high columns (originally fabricated by colloidal lithography ± inset, courtesy of Dr D. Sutherland iNANO, Denmark) imprinted into its leading edge. (b) Dynamin staining in similar fibroblasts on similar structures showing increased attempted endocytosis in lamellae membranes.

biomaterial substrate influence the adsorption of proteins from culture medium. This in turn can affect integrin-mediated cell adhesion [98]. The same cell type may attach to different substrates via different proteins. For example, one study demonstrated that mesenchymal stem cells attach to PCL primarily via vitronectin, whereas, on PLGA, type-I collagen mediates attachment [99]. Topography affects protein adsorption by causing local changes in surface chemistry and presenting areas that may give preferential adsorption. This may, in turn, provide more sites for focal adhesion formation. It has been proposed that in the case of grooves, adhesive proteins may co-localise with sharp discontinuities [100]. This mechanism for cell alignment has been disputed as it does not take into account the effect that different features' depths have on cell alignment or differences due to medium composition. Additionally, cells have been reported to align on surfaces without sharp features [101]. Oriented cells secrete similarly oriented extracellular matrix [100, 102], which in turn may lead to daughter cells adopting the same orientation. The conformation of adsorbed protein is affected by surface topography. Ridges with 1 nm height and 40 nm width on titanium have been reported to induce orientation of adsorbed F-actin, whereas similar structures with 4 nm heights, or flat titanium surfaces, resulted in lower adsorption of the same protein, which was also randomly oriented [103]. Germanium nanopyramids have also been used to study the effect of topography, with the authors excluding surface chemistry as a reason for the observed differences in protein adsorption [104]. Increased density of the nanopyramids (and thus increased surface roughness) resulted in a 2±3 fold increase in adsorption of bovine serum albumin and bovine -globulin compared to flat substrates. The globulin became inactive on the rougher surfaces, resulting in reduced release in vitro of interleukin-1 (one of the most crucial cytokines in inflammatory reactions) [105]. Furthermore, it has been demonstrated that adsorption of

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proteins onto nanorough surfaces leads to a masking of the surface features [106]. The combination of a pattern of surface chemistry and topography is also important; Denis et al. showed that the surface chemistry of a nanorough substrate affected the amount of protein adsorbed, whereas the surface topography modulated its organisation [106], although it has also been reported that initial protein adsorption is not affected by surface chemistry [107]. It must be noted that surface topography results in an increased surface area compared to a flat substrate, and therefore more protein may adsorb. Counterintuitively, however, it has been reported that when corrected for surface area, the amount of fibronectin adsorbed onto Ti surfaces with a micron-scale roughness was 50% less than that on smooth surfaces [108], although both surfaces promoted normal expression of two functional domains of the protein. A different study reports no difference in protein adsorption between nanorough and smooth Ti surfaces [31], although the authors did not account for differences in surface area. It appears that different surface topography affects protein adsorption in different ways, and the inconsistencies in reported results suggest that comprehensive understanding of this phenomenon is some way off.

7.4.4

Chemistry versus topography

In vivo structures, particularly basement membranes, have topography and chemistry, often grooves and ECM fibrils on the scale of 10±300 nm [109, 100]. While the focus of this chapter is on the cellular response to topographic features, it is also possible to elicit many similar responses by using chemical patterning (reviewed in [111]). In addition, many of the methods of producing topography result in deliberate or inadvertent chemical patterning, the effects of which have been briefly discussed in the section on protein adsorption. Equally, many techniques to create chemical patterns will also result in changes to surface topography, a factor which is rarely considered by investigators. A few studies have attempted to distinguish what happens when chemical and topographical patterns compete. There is no clear trend ± results appear to depend on the cell type [112±114]. In general, it has been proposed that chemical patterning regulates cell function and is scale-dependent, whereas the response to topographical patterning depends on whether it is isotropic or anisotropic [115].

7.5

Technologies and potential applications

7.5.1

Isolation and expansion of cells in vitro

The use of stem cells in tissue engineering not only opens up the potential to produce patient-specific tissues, reducing the risk of immune rejection, but through the understanding of material properties that elicit specific responses, could in the future allow the formation of complex tissues.

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Expansion of cells has traditionally incorporated serum or other animalderived products. As cell-based therapies move towards the clinic, there is a growing need for the use of animal-free systems to avoid the problems associated with this. Chemical modification of substrates, particularly using ECM components such as fibronectin and laminin has provided some success in maintaining pluripotency of ES and MSC [116, 117]. Other researchers have looked at using surface topography to improve cell expansion. The use of topographically patterned substrates for culturing cells has one clear advantage over the use of defined media ± it allows cell growth and development to be tailored to a specific application without the need to use potentially harmful chemicals in the body.

7.5.2

Embryonic stem cells

The main aims of biomaterials for stem cell culture are to provide cues that direct the differentiation of stem cells into a specific cell lineage and to allow the stem cells to remain undifferentiated but still undergo proliferation and self-renewal. Embryonic stem cell response to topography has been documented in terms of both controlled differentiation [118, 119] and retention of self-renewal/proliferative capabilities. On the latter topic, one such paper has looked at the effect of grooves with varying pitch from 400 nm up to the microscale at 4000 nm on human ES cell differentiation and self-renewal in the presence and absence of a feeder layer [120]. In culture, ES cells are prone to spontaneous differentiation, leading to a reduction in their capacity for self-renewal and pluripotentency. It has been shown that culture of human ES cells on nanogrooves in the presence of a feeder layer significantly reduces the rate of spontaneous differentiation relative to ES cells cultured on feeder layers alone [120]. Interestingly, however, this study also reported that when ES cells were cultured on the nanogrooves in the absence of a feeder layer, the rate of spontaneous differentiation increased compared with those cultured with only a feeder layer. These results indicate that ES cells may be influenced by topography in a contextual manner, and their response therefore also depends on other external factors. The ability to control self-renewal and proliferation of ES cells using topography has also been studied using mouse ES cells in response to threedimensional (3D) nanofibres, but importantly in the presence of only 5% of the original feeder layer [120, 121]. The study showed that mouse ES cells cultured on the 3D nanofibrillar surfaces had an increased rate of proliferation in comparison with those cultured solely on glass coverslips. The authors noted raised Nanog (a protein required for the maintenance of stem cell pluripotency) levels on the 3D nanofibres. These results indicate that the 3D nanofibre structures may provide an alternative method for maintaining self-renewal and proliferation capacities in mouse ES cells as effectively as feeder layers alone, eliminating the need for feeder layers or soluble factors.

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7.5.3

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Stem cell differentiation

Many soluble factors have been used to differentiate stem cells down different lineages; however, there is now compelling evidence that topography alone can produce the same effect.

7.5.4

Skeletal stem cells

The effect of nanotopographical variation on the differentiation of skeletal stem cells has been noted in several key studies. The first of these examined Stro-1 enriched skeletal stem cells cultured on nanopits (120 nm diameter, 100 nm depth) with varying degrees of disorder and geometry, ranging from an absolute square and controlled disorder, to a random arrangement embossed into the polymer polymethylmethacrylate (PMMA) [122]. The key finding from this study was that when cultured specifically on nanopits with a controlled disorder of 50 nm from their true centre (designated NSQ50), skeletal stem cells expressed particularly high levels of bone cell markers. This was comparable with the results from stem cells cultured on control planar substrates in medium supplemented with dexamethasone and ascorbic acid (chemical enhancers of skeletal stem cell differentiation down the osteogenic lineage) but in contrast to that of the square or random topographies, cells cultured on planar control substrates without osteogenic medium expressed negligible amounts of bone cell markers and appeared to have a bipolar fibroblastic-like appearance. Further evidence revealed an increase in expression of genes associated with bone cell development, comparable with those in cells on the flat surface in osteogenic medium, and considerably higher than expression levels from cells on the supplement-free planar controls. In a follow-up study, a temporal differentiation profile of skeletal stem cells cultured on the osteoinductive NSQ50 topography was carried out with reference to the classical osteogenic differentiation profile laid out by Stein and Lian [123]. This study used the skeletal stem cell markers Stro-1 and ALCAM (activated leukocyte cell adhesion molecule), together with the bone cell markers osteocalcin and osteopontin, to characterise the progression from an undifferentiated stem cell towards a committed bone cell. It was found that the skeletal stem cells cultured on the NSQ50 topography had a normal differentiation profile in line with the proposed osteogenic differentiation model. This provides further evidence that differentiation of skeletal stem cells cultured on nanotopography can produce an equal, if not superior, method for differentiation than chemical induction [74]. A recent study of the effect of carbon nanotube dimensions on skeletal stem cell fate revealed that by increasing the diameter size of the nanotubes from a range of 30 nm up to 100 nm, it was possible to alter the adhesion, elongation and differentiation of these stem cells [124]. It was shown that the 30 nm

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nanotubes had a higher number of adherent cells with a more rounded morphology, in contrast to that of the stem cells cultured on the 100 nm nanotubes, which developed highly elongated morphologies with a low level of cell adhesion. Osteogenic differentiation of the skeletal stem cells in this study was observed to occur on the carbon nanotubes with a 100 nm diameter, with negligible amounts of osteogenic markers observed on carbon nanotubes of 50 nm or less. Interestingly, the authors also investigated the relationship between pore size and initial cell density, relating to previously reported work [125] which indicated that a lower seeding density led to skeletal stem cell differentiation into osteoblasts, with a lower seeding density predisposing differentiation into adipocytes. The study also reported an inverse correlation between pore size and cell density on the 100 nm nanotubes, which ultimately led to an increase in osteogenic gene expression, consistent with other studies [72]. In a further study, skeletal stem cells were shown to differentiate down a nonskeletal lineage, a process known as transdifferentiation, and a somewhat controversial topic. The skeletal stem cells were seen to differentiate down a neuronal lineage as evidenced by the upregulation of mature neuronal cell markers, MAP2 (microtubule-associated protein 2) and -tubulin III, when cultured on nanogratings of 350 nm, in the absence of neuronal differentiation medium [126]. Expression of MAP2 was consistently higher in skeletal stem cells cultured over 14 days on the nanotopography without retinoic acid, a neuronal differentiation factor, than that of skeletal stem cells cultured on planar PDMS (poly(dimethylsiloxane)) with 1 M or 30 M retinoic acid, indicating nanotopography alone may have a stronger influence upon skeletal stem cell differentiation than chemical induction alone.

7.5.5

Neural stem cells

While a large body of work has investigated the effect of topography on skeletal stem cells, other stem cell sources have also been studied. A study using polymer films with a honeycomb-like microtopography indentified that these substrates have the ability to retain proliferation of neural stem cells without differentiation and provide cues for directed differentiation [127]. By varying the pore diameter of the honeycomb film it was possible to control whether rat neural stem cells differentiated into neurons, as indicated by expression of MAP2, or whether proliferation could occur without differentiation, as indicated by Nestin expression (a neuronal stem cell marker). It was found that by increasing the pore size of the honeycomb film above 5 m, the differentiation of neuronal stem cells into mature neurons was enhanced, with the effect being more prominent as the pore size was increased to 15 m. In contrast, below 5 m it was seen that differentiation was suppressed with the majority of cells expressing Nestin in the absence of MAP2. Neuronal stem cell differentiation has also been shown to occur on micropatterned grooved substrates [128] and electrospun fibres [129].

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For more details on differentiation of cells using biomaterial substrates, the reader is directed to [110, 130, 131].

7.6

Tissue regeneration

As has been mentioned previously, many natural tissues have topographical features. When trying to replicate this structure in vitro or in vivo, there is the potential to use surface topography to achieve the desired cellular behaviour and tissue architecture. There are many tissues in which this could be exploited. A few examples are given below.

7.6.1

Neural

Nerve regeneration is possible if the regrowing axons are provided with appropriate directional guidance. Much attention has been focussed on creating physical or chemical pathways for guiding axon regeneration. Often this is in the form of conduits that direct axon regrowth and prevent infiltration of other cell types. In terms of neural cell attachment, nanoscale roughness has been demonstrated to affect neuron cell attachment [132, 133]. Schwann cells may be able to mediate nerve regeneration in the peripheral and central nervous systems. They release neurotrophic factors and synthesise ECM, as well as expressing various cell adhesion molecules on their surface. Chemical guidance has also been used, both on its own [134] and in conjunction with topographical features [135]. Thompson and Buettner demonstrated Schwann cell alignment can direct neurite outgrowth in the absence of other topographical or chemical cues [136]. They speculated that the topography presented by the Schwann cells to the neurites provides contact guidance, in addition to molecular cues on the Schwann cell surface, possibly laminin. Neurons and Schwann cells cultured on PDMS with features designed to replicate the topography that Schwann cells provide were aligned to the surface features [137]. For an overview of various aspects of neural tissue engineering, including the use of surface topography, the reader is directed to [138] and Chapter 13.

7.6.2

Dental implantology

Whereas the use of a rough implant surface for osseointegration is well established (see [139] and Chapter 15), the dental implant-soft tissue may also benefit from surface topography. With a natural tooth, collagen fibres run laterally from the cementum through the soft tissue, perpendicular to the root, whereas with an implant the collagen fibres run parallel to the implant surface [140±142]. This parallel fibre orientation means that the connective tissue is only attached weakly to the implant and may be unable to prevent apical

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migration of epithelial tissue which may lead to gingival recession or bone resorption [143]. It has been speculated that the use of surface topography could be used to inhibit epithelial migration by preventing the cells from migrating and encouraging alignment of gingival fibroblasts. Oriented fibroblasts may secrete collagen fibres circumferentially around the abutment, strengthening the connection between the soft tissue and implant surface and preventing epithelial migration. Epithelial migration has been demonstrated to be inhibited by rough and micro-grooved surfaces in vitro [144±146]. In vitro studies have also established the ability of micron-scale grooves to induce orientation of gingival fibroblasts [61, 147, 148], which could produce oriented collagen fibres, enhancing tissue attachment, and restrict epithelial migration. This hypothesis has been tested in vivo, with studies reporting that micron-scale grooves, oriented perpendicular to the direction of migration of epithelial tissue, could inhibit epithelial downgrowth. The mechanism by which this occurred appeared to depend on the groove size; larger grooves (30 m wide, 22 m deep) enhanced the attachment and orientation of gingival fibroblasts, preventing epithelial migration, whereas smaller grooves (30 m wide, 10 m or 3 m deep) appeared to prevent the migration of epithelial cells by contact guidance [149]. A variety of rough abutment surfaces have also been reported to give a favourable connective tissue attachment compared to smooth surfaces [150, 151]. The occurrence of epithelial downgrowth is not limited to dental implantology ± it can occur in other percutaneous implants.

7.6.3

Artificial blood vessel

The successful engineering of artificial blood vessels is another area in which tissue engineering could provide a solution. Of particular importance is mimicking the structure of the natural vessel by production of circumferentially aligned ECM, which is reported to give the artery its natural mechanical properties and similarly aligned smooth muscle cells, which allow vessel contraction. Furthermore, cultured smooth muscle cells lose their contractility, as well as proliferating and synthesising ECM, which can lead to occlusion of the engineered vessel. It may be desirable, therefore, that the contractile phenotype is induced by the scaffold; microchannels have been reported to induce such a phenotype change as well as orienting the cells [152]. Cell-seeded collagen gel systems that aimed to recreate the laminate structure of the natural vessel lacked the mechanical properties required to withstand blood flow, possibly due to the lack of ECM and cell alignment [153]. Improved mechanical properties were obtained with a design where the collagen was in the form of aligned fibres, to which smooth muscle cells oriented [154], demonstrating the potential for exploitation of contact guidance in this application. Several studies report the ability of various polymeric scaffolds to induce orientation of smooth muscle cells in vitro [155±157]; however, to the best of our knowledge, there are no

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reports of similar response in vivo. Topography may provide cell alignment even when there is a competing orienting stimulus in the form of blood flow; it has been reported that endothelial cells remained aligned with micron-scale grooves even in the presence of relatively high fluid flow [158], although other studies report cell alignment in the direction of flow [159].

7.6.4

Contamination

Medical devices can be prone to infection, which can necessitate their removal. In the case of dentures or contact lenses this may be relatively mild, but surgery to excise a hip replacement following aseptic loosening is much more serious. Modifications to surface topography to elicit a particular response by the host tissue may result in inadvertently offering an opportunity for prokaryotic cells to thrive. Bacteria may come from a variety of sources, particularly the patient's skin or mucous membranes, and the hospital environment. Colonisation of the device surface can result in the formation of a bacterial biofilm that is resistance to host-defence mechanisms and antibiotic therapy. Detailed reviews of the mechanisms of bacterial adhesion to biomaterials are available [160, 161]. Many aspects of the biomaterial surface can affect bacterial adhesion; these include surface chemistry, wettability and surface roughness. Many studies have investigated the effect of micron-scale surface topography on microbial colonisation and in most cases, it is reported that rougher surface promote bacterial adhesion [162, 163]. A few reports suggest that surface roughness has no effect; it has been proposed that these differences are due to differences in the scale of surface roughness, surface chemistry and wettability and species of microorganism studied [164]. In particular, different microorganisms respond differently to features of different size; Whitehead et al. reported that whereas the retention of Staphylococcus aureus was greatest on surfaces with 2 m features, Candida albicans numbers were similar on surfaces with features ranging from 0.2 to 2 m [165]. Similar findings have been reported with different types of bacteria and algae [166, 167]. The `attachment point' theory [168, 169] suggests that if the fouling organism is larger than the scale of the topography, it will be weakly attached compared to an organism smaller than the surface features. This is due to the increase in hydrodynamic tension and surface area. In addition, surfaces features may afford protection from shear forces. The majority of bacteria are 0.5±10 m in size. It has been suggested that, broadly, a surface with Ra < 0.8 m is less likely to be colonised by bacteria [170]. In a similar way to eukaryotic cells, bacterial adhesion is affected by the shape of surface topography. A surface with features such as steep slopes and close packed peaks may prevent the cell conforming, thus reducing the number of contact points and weakening adhesion. Emerson et al. reported that the strength of Staphylococcus epidermis bonding on nanorough surfaces was weaker on rougher surfaces [171]. Another study reported that the attachment

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and growth of C. albicans was inhibited on surfaces coated with silica nanoparticles, speculating that the weak attachment may be due to modification in protein adsorption or the dissolution of the particles [172]. These studies show that the effect of nanoscale topography on microbial adhesion is often different from that observed with micron-scale topography. The effect of surface finish of some medical devices on the potential for bacterial adhesion has been investigated. In general, smooth surfaces are reported to be less susceptible to contamination [173], although several studies relating to medical devices suggest that there is a threshold roughness (Ra < 0.2 m) [162, 174, 175] below which further improvements do not occur. The effect of the surface roughness of the abutment of a dental implant on plaque bacteria adhesion has been studied in vivo and results do not always support in vitro findings: one study in humans reported no difference in dental plaque accumulation on surfaces with different roughness [176]. The roughness values were around Sa = 1 m, similar to commercially available implants, although the study was over a relatively short time period in healthy volunteers. The surfaces of central venous catheters, intraocular lenses and biliary stents are also reported to be more likely to be colonised by bacteria if they have rough surfaces [163, 177, 178]. In many cases, the potential for contamination can be minimised by effective sterilisation and handling. However, with some devices, particularly those that are percutaneous or are used outside of the body, such as contact lenses and dentures, contamination is a constant threat and the potential of surface modification to enhance or minimise the risk should be considered in the design of such devices.

7.6.5

Inflammatory response

In addition to an increased potential for microbial contamination, the surface roughness of a material may affect the inflammatory response. Several studies have investigated the influence of surface roughness on inflammation and a review of how surfaces affect macrophages is available [179]. In vitro, it has been demonstrated that a variety of cell types involved in the inflammatory response, including macrophages, leukocytes and granulocytes accumulate on rough surfaces to a greater extent than smooth surfaces [180±183]. In addition to increased attachment, macrophages are reported to secrete more proinflammatory cytokines on rough titanium [184] substrates. In vivo studies also indicate that rough surfaces are associated with an increased number of inflammatory cells [185, 186], although interestingly a study by Parker et al. [187] reported an increased number of inflammatory cells at the implant±tissue interface of a randomly rough surface but not for a microgrooved surface. It has been proposed that surface topography may mediate the inflammatory response due to differences in protein adsorption [188]. Koh et al. investigated the effect of PLGA-carbon nanotube composites on platelet adhesion [189]. They found

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that platelet adhesion was influenced by surface topography, and that this correlated with fibrinogen adsorption. They also suggested that fibrinogen adopted different conformations depending on the surface chemistry and topography, indicating that consideration of the effect of any surface modification on protein adsorption may be useful in improving the design of medical devices in respect of the inflammatory response they may elicit. Blood-contacting devices may induce an adverse inflammatory response due to the potential of biomaterials to activate the complement cascade. Complement activation is affected by chemical groups present on biomaterial substrates. Hydroxyl groups, for example, are strong activators [190, 191], whereas amino groups activate the cascade to a lesser extent [192]. In the case of activation by amino groups, it was suggested that this is due to altered protein adsorption onto the material surface. Although little work is available on the effect of surface topography on complement activation, it is possible to speculate that this type of surface modification could result in complement activation due to particular patterns of protein adsorption.

7.7

Current issues and future trends

The majority of research into the use of topography to control cellular response has been carried out in the laboratory. It is desirable to apply this knowledge to clinical situations. Central to aims of moving to clinic will be firstly, the development of biomaterials with simple and then complex designs and, secondly, development into 3D for tissue engineering applications. Many of the techniques used to create topography are 2D; this is especially true for the most high-resolution methodologies such as electron beam lithography. Continual development in silicon-based technologies will push this technique to create smaller features over larger areas, but the nature of the serial electron exposure means the fabrication needs to be performed on flat substrates. Other techniques, such as phase separation, etc., require the rapid drying of a thin film to allow demixing, again, this presents problems for the move to 3D. A first step will be to produce medical devices that are flat, gently curved, or that have planar areas (e.g. implant surfaces). Even here there are significant hurdles ± these are now being overcome. When considering scale-up, it would be impossible to conceive the individual fabrication of each medical device; it would cost too much in terms of time, money and development of production line. However, rapid reproduction into thermopolymers is now possible down to 5 nm resolution [11]. On DVD production lines, a master `shim' is used to injection mould subsequent `copies' in polycarbonate and this is what we play in DVD players ± quick, cheap and simple. This can also be performed with polymers such as polymethylemethacrylate, polylactide and polycaprolactone. It could be envisaged that nanopatterned 2D polymer substrates could be used, for example, for maxillofacial surgery or in orthopaedics to help slow healing

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fractures. However, a major problem for orthopaedic application is that thermopolymers are not suitable load-bearing materials. There are two possible approaches to this problem, firstly the use of composite materials and then, secondly, embossing into metals. The first approach is based on compositing a mechanically suitable filler into a polymer, e.g. hydroxyapatite (HA) into polylactic acid (PLA) or high-density polyethylene (HDPE). HA has a high Young's modulus but is brittle, PLA and HDPE have low Young's modulus and fail in a ductile manner [193]. The combination produces something closer to cortical bone that has a surface that can be embossed into at glass transition temperature of the polymer. The second route requires the production of very hard (e.g. diamond) shims or use of, for example, nano-imprint lithography (NIL [194]) and a wet etch to transfer patterns into metals. It could be envisaged that this could be done on flat side panels of the implants leading to improved secondary fixation. For the transition to 3D an exciting recent advance has been the development of the topographical `Swiss-roll' [195]. Here, a biodegradable polymer is imprinted on both sides (i.e. between two shims). The nanopattern is imprinted on one side and, on the other, large (tens of micron) `spacers' are embossed into the material. The sheet is then rolled with the spacers acting to allow media and oxygen perfusion to the core of the implant and to allow cell contact with the nanopatterns. It could be envisaged that these rolls pre-seeded with cells in the lab could be used to fill larger defects. In addition to bone, these rolls are being developed as capillary and neuronal conduits by varying the imprinted patterns to suit different cells. In addition to development of fabrication techniques, it is yet to be determined whether the results from highly controlled in vitro studies are transferrable to the complex in vivo environment. There are many topographical and chemical stimuli that may compete with the topography provided by a biomaterial substrate. However, work by Engler et al. suggests that the cues from the substrate are powerful and may be able to override e.g. soluble factors [196]. This may depend, in part, on cell density ± i.e., individual cells may be able to gain a lot of information from the materials, but large groups of cells will perceive far more, e.g. cytokine and cadherin signalling. The local environment to the cell must be critical to function. It is possible to speculate that this is exemplified in embryogenesis. The cells of the blastocyst are identical ± similar DNA, similar transcription factors, etc. However, chemical diffusion into the blastocyst, the change in stiffness in different regions and the change in curvature present chemical and topographical cues and also local changes in Young's modulus. These factors, including topography, may well trigger the formation of the different germ lines. Materials scientists need to take these powerful cues and make materials designed around the cell. This can only be achieved by far more basic research.

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7.8

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Acknowledgements

VRK is supported by a private local charity. MJD and RJM are supported by grants from BBSRC, EPSRC, Chief Scientist Office (NHS) and University of Glasgow.

7.9

References

1. Carrel, A. and M. Burrows, Culture in vitro of malignant tumors. J Exp Med, 1911. 12: 571±5. 2. Curtis, A.S.G. and M. Varde, Control of cell behaviour: topological factors. J Nat Cancer Res Inst, 1964. 33: 15±26. 3. Wilkinson, C.D.W., et al., The use of materials patterned on a nano- and micrometric scale in cellular engineering. Mat Sci Eng, 2002. 19: 263±9. 4. Dalby, M.J., et al., Investigating the limits of filopodial sensing: a brief report using SEM to image the interaction between 10 nm high nano-topography and fibroblast filopodia. Cell Biol Int, 2004. 28(3): 229±36. 5. Clark, P., et al., Topographical control of cell behaviour. I. Simple step cues. Development, 1987. 99(3): 439±48. 6. Clark, P., et al., Topographical control of cell behaviour: II. Multiple grooved substrata. Development, 1990. 108(4): 635±44. 7. Dunn, G.A. and A.F. Brown, Alignment of fibroblasts on grooved surfaces described by a simple geometric transformation. J Cell Sci, 1986. 83: 313±40. 8. Clark, P., et al., Cell guidance by ultrafine topography in vitro. J Cell Sci, 1991. 99 (Pt 1): 73±7. 9. Moore, G.E., Cramming more components onto integrated circuits. Electronics, 1965. 38(8): 114±17. 10. Gallagher, J.O., et al., Interaction of animal cells with ordered nanotopography. IEEE Transactions on Nanobioscience, 2002. 1(1): 24±8. 11. Gadegaard, N., S. Mosler, and N.B. Larsen, Biomimetic polymer nanostructures by injection moulding. Macromol Mat Eng, 2003. 288: 76±83. 12. Vieu, C., et al., Electron beam lithography: resolution limits and applications. Appl Surf Sci, 2000. 164: 111±17. 13. Denis, F.A., et al., Fabrication of nanostructured polymer surfaces using colloidal lithography and spin coating. Nanoletters, 2002. 2: 1419±25. 14. Hanarp, P., et al., Control of nanoparticle film structure for colloidal lithography. Colloids Surfaces A, 2003. 214: 23±36. 15. Affrossman, S., et al., Surface topography and composition of deuterated polystyrene±poly(bromostyrene) blends. Macromolecules, 1996. 29(14): 5010± 5016. 16. Affrossman, S., S.A. O'Neill, and M. Stamm, Topography and surface composition of thin films of blends of polystyrene with brominated polystyrenes: effects of varying the degree of bromination and annealing. Macromolecules, 1998. 31: 6280± 8. 17. Affrossman, S., et al., Surface structure of thin film blends of polystyrene and poly(n-butyl methacrylate). Colloid Polym Sci, 2000. 278: 993±9. 18. Affrossman, S. and M. Stamm, The effect of molecular weight on the topography of thin films of blends of poly(4-bromostyrene) and polystyrene. Colloid Polym Sci, 2000. 278: 888±93.

ß Woodhead Publishing Limited, 2011

Biomaterial surface topography to control cellular response

193

19. Krishnamoorthy, S., et al., Block copolymer micelles as switchable templates for nanofabrication. Langmuir, 2006. 22(8): 3450±2. 20. Mustafa, K., et al., Attachment and proliferation of human oral fibroblasts to titanium surfaces blasted with TiO2 particles. A scanning electron microscopic and histomorphometric analysis. Clin Oral Implants Res, 1998. 9(3): 195±207. 21. Mwenifumbo, S., et al., Cell/surface interactions on laser micro-textured titaniumcoated silicon surfaces. J Mat Sci: Mat Med, 2007. 18(1): 9±23. 22. Wojciak-Stothard, B., et al., Guidance and activation of murine macrophages by nanometric scale topography. Exp Cell Res, 1996. 223(2): 426±35. 23. Manwaring, M.E., J.F. Walsh, and P.A. Tresco, Contact guidance induced organization of extracellular matrix. Biomaterials, 2004. 25(17): 3631±8. 24. Teixeira, A.I., et al., Epithelial contact guidance on well-defined micro- and nanostructured substrates. J Cell Sci, 2003. 116(10): 1881±92. 25. Fraser, S.A., et al., Sub-micron and nanoscale feature depth modulates alignment of stromal fibroblasts and corneal epithelial cells in serum-rich and serum-free media. J Biomed Mat Res Part A, 2008. 86A(3): 725±35. 26. den Braber, E.T., et al., Quantitative analysis of cell proliferation and orientation on substrata with uniform parallel surface micro-grooves. Biomaterials, 1996. 17(11): 1093±9. 27. Crouch, A.S., et al., Correlation of anisotropic cell behaviors with topographic aspect ratio. Biomaterials, 2009. 30(8): 1560±7. 28. Deligianni, D.D., et al., Effect of surface roughness of hydroxyapatite on human bone marrow cell adhesion, proliferation, differentiation and detachment strength. Biomaterials, 2001. 22(1): 87±96. 29. Green, A.M., et al., Fibroblast response to microtextured silicone surfaces: Texture orientation into or out of the surface. J Biomed Mat Res, 1994. 28(5): 647±53. 30. Mata, A., et al., Growth of connective tissue progenitor cells on microtextured polydimethylsiloxane surfaces. J Biomed Mat Res, 2002. 62(4): 499±506. 31. Cai, K., J. Bossert, and K.D. Jandt, Does the nanometre scale topography of titanium influence protein adsorption and cell proliferation? Colloids and Surfaces B: Biointerfaces, 2006. 49(2): 136±144. 32. Martin, J.Y., et al., Effect of titanium surface roughness on proliferation, differentiation, and protein synthesis of human osteoblast-like cells (MG63). J Biomed Mat Res, 1995. 29(3): 389±401. 33. Wan, Y., et al., Adhesion and proliferation of OCT-1 osteoblast-like cells on microand nano-scale topography structured poly(L-lactide). Biomaterials, 2005. 26(21): 4453±9. 34. Zinger, O., et al., Time-dependent morphology and adhesion of osteoblastic cells on titanium model surfaces featuring scale-resolved topography. Biomaterials, 2004. 25(14): 2695±711. 35. Fan, Y.W., et al., Culture of neural cells on silicon wafers with nano-scale surface topograph. J Neurosci Meth, 2002. 120(1): 17±23. 36. Dalby, M.J., D. Pasqui, and S. Affrossman, Cell response to nano-islands produced by polymer demixing: a brief review. IEE Proc Nanobiotechnol, 2004. 151: 53±61. 37. Dalby, M.J., et al., In vitro reaction of endothelial cells to polymer demixed nanotopography. Biomaterials, 2002. 23(14): 2945±54. 38. Dalby, M.J., et al., Polymer-demixed nanotopography: control of fibroblast spreading and proliferation. Tissue Eng, 2002. 8(6): 1099±108. 39. Dalby, M.J., et al., Osteoprogenitor response to semi-ordered and random nanotopographies. Biomaterials, 2006. 27(15): 2980±7.

ß Woodhead Publishing Limited, 2011

194

Surface modification of biomaterials

40. Dalby, M.J., et al., Nonadhesive nanotopography: fibroblast response to poly(nbutyl methacrylate)-poly(styrene) demixed surface features. J Biomed Mater Res, 2003. 67A(3): 1025±32. 41. Dalby, M.J., et al., Changes in fibroblast morphology in response to nano-columns produced by colloidal lithography. Biomaterials, 2004. 25(23): 5415±22. 42. Dalby, M.J., et al., Fibroblast response to a controlled nanoenvironment produced by colloidal lithography. J Biomed Mater Res, 2004. 69A(2): 314±22. 43. Andersson, A.S., et al., Nanoscale features influence epithelial cell morphology and cytokine production. Biomaterials, 2003. 24(20): 3427±36. 44. Andersson, A.S., et al., Influence of systematically varied nanoscale topography on the morphology of epithelial cells. IEEE Transactions on Nanobioscience, 2003. 2(2): 49±57. 45. Rice, J.M., et al., Quantitative assessment of the response of primary derived human osteoblasts and macrophages to a range of nanotopography surfaces in a single culture model in vitro. Biomaterials, 2003. 24(26): 4799±818. 46. Choi, C.H., et al., Cell interaction with three-dimensional sharp-tip nanotopography. Biomaterials, 2007. 28(9): 1672±9. 47. Dalby, M.J., et al., Investigating filopodia sensing using arrays of defined nano-pits down to 35 nm diameter in size. Int J Biochem Cell Biol, 2004. 36(10): 2015±25. 48. Dalby, M.J., et al., Osteoprogenitor response to defined topographies with nanoscale depths. Biomaterials, 2006. 27: 1306±15. 49. Curtis, A.S., et al., Cells react to nanoscale order and symmetry in their surroundings. IEEE Trans Nanobioscience, 2004. 3(1): 61±5. 50. Hart, A., et al., Osteoprogenitor response to low-adhesion nanotopographies originally fabricated by electron beam lithography. J Mater Sci Mater Med, 2007. 18(6): 1211±18. 51. Dalby, M.J., et al., Group analysis of regulation of fibroblast genome on lowadhesion nanostructures. Biomaterials, 2007. 28(10): 1761±9. 52. Dalby, M.J., N. Gadegaard, and C.D.W. Wilkinson, The response of fibroblasts to hexagonal nanotopography fabricated by electron beam lithography. J Biomed Mat Res ± Part A, 2008. 84(4): 973±9. 53. Dalby, M.J., Cellular response to low adhesion nanotopographies. Int J Nanomed, 2007. 2(3): 373±81. 54. Ohara, P.T. and R.C. Buck, Contact guidance in vitro: a light, transmission, and scanning electron microscopic study. Exp Cell Res, 1979. 121(2): 235±49. 55. Clark, P., et al., Topographical control of cell behaviour. I. Simple step cues. Development, 1987. 99(3): 439±48. 56. Clark, P., et al., Topographical control of cell behaviour: II. Multiple grooved substrata. Development, 1990. 108(4): 635±44. 57. Oakley, C. and D.M. Brunette, The sequence of alignment of microtubules, focal contacts and actin filaments in fibroblasts spreading on smooth and grooved titanium substrata. J Cell Sci, 1993. 106(1): 343±54. 58. Brunette, D.M., Fibroblasts on micromachined substrata orient hierarchically to grooves of different dimensions. Exp Cell Res, 1986. 164(1): 11±26. 59. den Braber, E.T., et al., Quantitative analysis of fibroblast morphology on microgrooved surfaces with various groove and ridge dimensions. Biomaterials, 1996. 17(21): 2037±44. 60. Curtis, A. and C. Wilkinson, Topographical control of cells. Biomaterials, 1997. 18(24): 1573±83. 61. Oakley, C., N.A. Jaeger, and D.M. Brunette, Sensitivity of fibroblasts and their

ß Woodhead Publishing Limited, 2011

Biomaterial surface topography to control cellular response

62. 63. 64. 65. 66. 67. 68. 69. 70. 71. 72. 73. 74. 75. 76. 77. 78. 79. 80. 81.

195

cytoskeletons to substratum topographies: topographic guidance and topographic compensation by micromachined grooves of different dimensions. Exp Cell Res, 1997. 234(2): 413±24. Hamilton, D.W. and D.M. Brunette, `Gap guidance' of fibroblasts and epithelial cells by discontinuous edged surfaces. Exp Cell Res, 2005. 309(2): 429±37. Dunn, G.A., How do cells respond to ultrafine surface contours? BioEssays, 1991. 13(10): 541±3. Teixeira, A.I., et al., The effect of environmental factors on the response of human corneal epithelial cells to nanoscale substrate topography. Biomaterials, 2006. 27(21): 3945±54. Berry, C.C., et al., The influence of microscale topography on fibroblast attachment and motility. Biomaterials, 2004. 25(26): 5781±8. Dalby, M.J., et al., Investigating the limits of filopodial sensing: a brief report using SEM to image the interaction between 10 nm high nano-topography and fibroblast filopodia. Cell Biol Int, 2004. 28(3): 229±36. Dalby, M.J., et al., Nanotopographical stimulation of mechanotransduction and changes in interphase centromere positioning. J Cell Biochem, 2007. 100(2): 326±38. Clark, E.A. and J.S. Brugge, Integrins and signal transduction pathways: the road taken. Science, 1995. 268(5208): 233±9. Chen, C.S., et al., Geometric control of cell life and death. Science, 1997. 276(5317): 1425±8. Hall, A., Rho GTPases and the actin cytoskeleton. Science, 1998. 279(5350): 509± 14. Sordella, R., et al., Modulation of Rho GTPase signaling regulates a switch between adipogenesis and myogenesis. Cell, 2003. 113(2): 147±58. McBeath, R., et al., Cell shape, cytoskeletal tension, and RhoA regulate stem cell lineage commitment. Developmental Cell, 2004. 6(4): 483±95. Hamilton, D.W. and D.M. Brunette, The effect of substratum topography on osteoblast adhesion mediated signal transduction and phosphorylation. Biomaterials, 2007. 28(10): 1806±19. Kantawong, F., et al., Whole proteome analysis of osteoprogenitor differentiation induced by disordered nanotopography and mediated by ERK signalling. Biomaterials, 2009. 30(27): 4723±31. Biggs, M.J., et al., Interactions with nanoscale topography: adhesion quantification and signal transduction in cells of osteogenic and multipotent lineage. J Biomed Mater Res A, 2008. 91a(1): 195±208. Biggs, M.J., et al., The use of nanoscale topography to modulate the dynamics of adhesion formation in primary osteoblasts and ERK/MAPK signalling in STRO-1+ enriched skeletal stem cells. Biomaterials, 2009. 30(28): 5094±103. Biggs, M.J., et al., Adhesion formation of primary human osteoblasts and the functional response of mesenchymal stem cells to 330nm deep microgrooves. J R Soc Interface, 2008. 5(27) 1231±42. Dalby, M.J., Topographically induced direct cell mechanotransduction. Med Eng Phys, 2005. 27(9): 730±42. Ingber, D.E., Tensegrity II. How structural networks influence cellular information processing networks. J Cell Sci, 2003. 116(Pt 8): 1397±408. Ingber, D.E., Tensegrity I. Cell structure and hierarchical systems biology. J Cell Sci, 2003. 116(Pt 7): 1157±73. Ingber, D.E., Mechanosensation through integrins: cells act locally but think globally. Proc Natl Acad Sci USA, 2003. 100(4): 1472±4.

ß Woodhead Publishing Limited, 2011

196

Surface modification of biomaterials

82. Maniotis, A.J., C.S. Chen, and D.E. Ingber, Demonstration of mechanical connections between integrins, cytoskeletal filaments, and nucleoplasm that stabilize nuclear structure. Proc Natl Acad Sci USA, 1997. 94(3): 849±54. 83. Maniotis, A.J., K. Bojanowski, and D.E. Ingber, Mechanical continuity and reversible chromosome disassembly within intact genomes removed from living cells. J Cell Biochem, 1997. 65(1): 114±30. 84. Ingber, D.E., Cellular tensegrity: defining new rules of biological design that govern the cytoskeleton. J Cell Sci, 1993. 104 (Pt 3): 613±27. 85. Charras, G.T. and M.A. Horton, Single cell mechanotransduction and its modulation analyzed by atomic force microscope indentation. Biophys J, 2002. 82(6): 2970±81. 86. Dalby, M.J., et al., Use of nanotopography to study mechanotransduction in fibroblasts±methods and perspectives. Eur J Cell Biol, 2004. 83(4): 159±69. 87. Heslop-Harrison, J.S., Comparative genome organization in plants: from sequence and markers to chromatin and chromosomes. Plant Cell, 2000. 12(5): 617±36. 88. Heslop-Harrison, J.S., A.R. Leitch, and T. Schwarzacher, The physical organisation of interphase nuclei, in The Chromosome, Ed. by J.S. Heslop-Harrison and R.B. Flavell. 1993, Bios Scientific Publishers, Oxford, pp. 221±32. 89. Mosgoller, W., et al., Chromosome arrangements in human fibroblasts at mitosis. Hum Genet, 1991. 88(1): 27±33. 90. Wang, N. and Z. Suo, Long-distance propagation of forces in a cell. Biochem Biophys Res Commun, 2005. 328(4): 1133±8. 91. Foster, H.A. and J.M. Bridger, The genome and the nucleus: a marriage made by evolution. Genome organisation and nuclear architecture. Chromosoma, 2005. 114(4): 212±29. 92. Bustamante, C., Z. Bryant, and S.B. Smith, Ten years of tension: single-molecule DNA mechanics. Nature, 2003. 421(6921): 423±7. 93. Osborne, C.S., et al., Active genes dynamically colocalize to shared sites of ongoing transcription. Nat Genet, 2004. 36(10): 1065±71. 94. Curtis, A.S.G., M.J. Dalby, and N. Gadegaard, Cell signaling arising from nanotopography: implications for nanomedical devices. Nanomedicine, 2006. 1(1): 67±72. 95. Curtis, A.S.G., M.J. Dalby, and N. Gadegaard, Nanoimprinting onto cells. J Roy Soc Interface, 2006. 3: 393±8. 96. Wood, M.A., Bagnaninchi, and M.J. Dalby, The beta integrins and cytoskeletal nanoimprinting. Exp Cell Res, 2008. 314(4): 927±35. 97. Dalby, M.J., et al., Attempted endocytosis of nano-environment produced by colloidal lithography by human fibroblasts. Exp Cell Res, 2004. 295(2): 387±94. 98. Keselowsky, B.G., D.M. Collard, and A.J. Garcia, Surface chemistry modulates focal adhesion composition and signaling through changes in integrin binding. Biomaterials, 2004. 25(28): 5947±54. 99. Chastain, S.R., et al., Adhesion of mesenchymal stem cells to polymer scaffolds occurs via distinct ECM ligands and controls their osteogenic differentiation. J Biomed Mat Res Part A, 2006. 78A(1): 73±85. 100. den Braber, E.T., et al., Orientation of ECM protein deposition, fibroblast cytoskeleton, and attachment complex components on silicone microgrooved surfaces. J Biomed Mat Res, 1998. 40(2): 291±300. 101. Walboomers, X.F., et al., Contact guidance of rat fibroblasts on various implant materials. J Biomed Mat Res, 1999. 47(2): 204±12. 102. Wang, J.H., et al., Cell orientation determines the alignment of cell-produced

ß Woodhead Publishing Limited, 2011

Biomaterial surface topography to control cellular response

197

collagenous matrix. J Biomech, 2003. 36(1): 97±102. 103. Galli, C., et al., Protein adsorption on topographically nanostructured titanium. Surf Sci, 2001. 474(1±3): L180±4. 104. Muller, B., et al., Impact of nanometer-scale roughness on contact-angle hysteresis and globulin adsorption. J Vacuum Sci Technol B, 2001. 19(5): 1715±20. 105. Riedel, M., B. MuÈller, and E. Wintermantel, Protein adsorption and monocyte activation on germanium nanopyramids. Biomaterials, 2001. 22(16): 2307±16. 106. Denis, F.A., et al., Protein adsorption on model surfaces with controlled nanotopography and chemistry. Langmuir, 2002. 18(3): 819±28. 107. Sutherland, D.S., et al., Influence of nanoscale surface topography and chemistry on the functional behaviour of an adsorbed model macromolecule. Macromol Biosci, 2001. 1(6): 270±3. 108. Francois, P., et al., Influence of surface treatments developed for oral implants on the physical and biological properties of titanium. (II) Adsorption isotherms and biological activity of immobilized fibronectin. Clin Oral Implants Res, 1997. 8(3): 217±25. 109. Abrams, G.A., et al., Nanoscale topography of the corneal epithelial basement membrane and Descemet's membrane of the human. Cornea, 2000. 19(1): 57±64. 110. Chai, C. and K.W. Leong, Biomaterials approach to expand and direct differentiation of stem cells. Mol Ther, 2007. 15(3): 467±80. 111. Hasirci, V. and H. Kenar. Novel surface patterning approaches for tissue engineering and their effect on cell behavior. 2006. Available from: http:// www.futuremedicine.com.ezproxy.liv.ac.uk/doi/abs/10.2217/17435889.1.1.73. 112. Britland, S., et al., Synergistic and hierarchical adhesive and topographic guidance of BHK cells. Exp Cell Res, 1996. 228(2): 313±25. 113. Charest, J.L., et al., Combined microscale mechanical topography and chemical patterns on polymer cell culture substrates. Biomaterials, 2006. 27(11): 2487±94. 114. Gomez, N., et al., Immobilized nerve growth factor and microtopography have distinct effects on polarization versus axon elongation in hippocampal cells in culture. Biomaterials, 2007. 28(2): 271±84. 115. Lim, J.Y. and H.J. Donahue, Cell sensing and response to micro- and nanostructured surfaces produced by chemical and topographic patterning. Tissue Engineering, 2007. 13(8): 1879±91. 116. Amit, M., et al., Feeder layer- and serum-free culture of human embryonic stem cells. Biol Reprod, 2004. 70(3): 837±45. 117. Matsubara, T., et al., A new technique to expand human mesenchymal stem cells using basement membrane extracellular matrix. Biochem Biophys Res Com, 2004. 313(3): 503±8. 118. Levenberg, S., et al., Differentiation of human embryonic stem cells on threedimensional polymer scaffolds. Proc Natl Acad Sci USA, 2003. 100(22): 12741±6. 119. Smith, L.A., et al., The influence of three-dimensional nanofibrous scaffolds on the osteogenic differentiation of embryonic stem cells. Biomaterials, 2009. 30(13): 2516±22. 120. McFarlin, D.F., K.J. Finn, P.F. Nealey, and C.J. Murphy, Nanoscale topography modulates human embryonic stem cell self-renewal. The American Society for Cell Biology 46th Annual Meeting, 2006. 121. Nur, E.K.A., et al., Three-dimensional nanofibrillar surfaces promote self-renewal in mouse embryonic stem cells. Stem Cells, 2006. 24(2): 426±33. 122. Dalby, M.J., et al., The control of human mesenchymal cell differentiation using nanoscale symmetry and disorder. Nat Mater, 2007. 6(12): 997±1003.

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

123. Stein, G.S. and J.B. Lian, Molecular mechanisms mediating proliferation/ differentiation interrelationships during progressive development of the osteoblast phenotype. Endocr Rev, 1993. 14(4): 424±42. 124. Oh, S., et al., Stem cell fate dictated solely by altered nanotube dimension. Proc Natl Acad Sci, 2009. 106(7): 2130±5. 125. Pittenger, M.F., et al., Multilineage potential of adult human mesenchymal stem cells. Science, 1999. 284(5411): 143±7. 126. Yim, E.K., S.W. Pang, and K.W. Leong, Synthetic nanostructures inducing differentiation of human mesenchymal stem cells into neuronal lineage. Exp Cell Res, 2007. 313(9): 1820±9. 127. Tsuruma, A., et al., Control of neural stem cell differentiation on honeycomb films. Colloids Surfaces A: Physicochem Eng Aspects, 2008. 313±314: 536±40. 128. Recknor, J.B., D.S. Sakaguchi, and S.K. Mallapragada, Directed growth and selective differentiation of neural progenitor cells on micropatterned polymer substrates. Biomaterials, 2006. 27(22): 4098±108. 129. Christopherson, G.T., H. Song, and H.Q. Mao, The influence of fiber diameter of electrospun substrates on neural stem cell differentiation and proliferation. Biomaterials, 2009. 30(4): 556±64. 130. Dawson, E., et al., Biomaterials for stem cell differentiation. Adv Drug Del Rev, 2008. 60(2): 215±28. 131. MartõÂnez, E., et al., Stem cell differentiation by functionalized micro- and nanostructured surfaces. Nanomedicine, 2009. 4(1): 65±82. 132. Cyster, L.A., et al., The effect of surface chemistry and nanotopography of titanium nitride (TiN) films on primary hippocampal neurones. Biomaterials, 2004. 25(1): 97±107. 133. Fan, Y.W., et al., Adhesion of neural cells on silicon wafer with nano-topographic surface. Appl Surf Sci, 2002. 187(3±4): 313. 134. Song, H.K., et al., Micropatterns of positive guidance cues anchored to polypyrrole doped with polyglutamic acid: a new platform for characterizing neurite extension in complex environments. Biomaterials, 2006. 27(3): 473±84. 135. Miller, C., S. Jeftinija, and S. Mallapragada, Micropatterned Schwann cell-seeded biodegradable polymer substrates significantly enhance neurite alignment and outgrowth. Tissue Engineering, 2001. 7(6): 705±15. 136. Thompson, D. and H. Buettner, Neurite outgrowth is directed by Schwann cell alignment in the absence of other guidance cues. Ann Biomed Eng, 2006. 34(1): 161±8. 137. Bruder, J.M., A.P. Lee, and D. Hoffman-Kim, Biomimetic materials replicating Schwann cell topography enhance neuronal adhesion and neurite alignment in vitro. J Biomat Sci, Polym Ed, 2007. 18: 967±82. 138. Schmidt, C.E. and J.B. Leach, Neural tissue engineering: strategies for repair and regeneration. Ann Rev Biomed Eng, 2003. 5(1): 293±347. 139. Stanford, C.M., Surface modifications of dental implants. Australian Dental J, 2008. 53(s1): S26±S33. 140. Abrahamsson, I., et al., The peri-implant hard and soft tissues at different implant systems. A comparative study in the dog. Clin Oral Implants Res, 1996. 7(3): 212±19. 141. Abrahamsson, I., et al., The mucosal attachment to titanium implants with different surface characteristics: an experimental study in dogs. J Clin Periodontol, 2002. 29(5): 448±55. 142. Comut, A.A., et al., Connective tissue orientation around dental implants in a canine model. Clin Oral Implants Res, 2001. 12(5): 433±40.

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143. Rompen, E., et al., The effect of material characteristics, of surface topography and of implant components and connections on soft tissue integration: a literature review. Clin Oral Implants Res, 2006. 17: 55±67. 144. Dalton, B.A., et al., Modulation of epithelial tissue and cell migration by microgrooves. J Biomed Mat Res, 2001. 56(2): 195±207. 145. Diehl, K.A., et al., Nanoscale topography modulates corneal epithelial cell migration. J Biomed Mat Res Part A, 2005. 75A(3): 603±11. 146. Fitton, J.H., et al., Surface topography can interfere with epithelial tissue migration. J Biomed Mat Res, 1998. 42(2): 245±57. 147. Inoue, T., et al., Effect of the surface geometry of smooth and porous-coated titanium alloy on the orientation of fibroblasts in vitro. J Biomed Mat Res, 1987. 21(1): 107±26. 148. KoÈunoÈnen, M., et al., Effect of surface processing on the attachment, orientation, and proliferation of human gingival fibroblasts on titanium. J Biomed Mat Res, 1992. 26(10): 1325±41. 149. Chehroudi, B., T.R. Gould, and D.M. Brunette, Titanium-coated micromachined grooves of different dimensions affect epithelial and connective-tissue cells differently in vivo. J Biomed Mat Res, 1990. 24(9): 1203±19. 150. Kim, H., et al., Effects of surface topography on the connective tissue attachment to subcutaneous implants. Int J Oral Maxillofacial Implants, 2006. 21(3): 354±65. 151. Squier, C.A. and P. Collins, The relationship between soft-tissue attachment, epithelial downgrowth and surface porosity. J Periodontal Res, 1981. 16(4): 434±40. 152. Shen, J.Y., et al., Three-dimensional microchannels in biodegradable polymeric films for control orientation and phenotype of vascular smooth muscle cells. Tissue Engineering, 2006. 12(8): 2229±40. 153. Weinberg, C.B. and E. Bell, A blood vessel model constructed from collagen and cultured vascular cells. Science, 1986. 231(4736): 397±400. 154. Tranquillo, R.T., et al., Magnetically orientated tissue-equivalent tubes: application to a circumferentially orientated media-equivalent. Biomaterials, 1996. 17(3): 349± 57. 155. In Jeong, S., et al., Tissue-engineered vascular grafts composed of marine collagen and PLGA fibers using pulsatile perfusion bioreactors. Biomaterials, 2007. 28(6): 1115±22. 156. Xu, C.Y., et al., Aligned biodegradable nanofibrous structure: a potential scaffold for blood vessel engineering. Biomaterials, 2004. 25(5): 877±86. 157. Zorlutuna, P., A. Elsheikh, and V. Hasirci, Nanopatterning of collagen scaffolds improve the mechanical properties of tissue engineered vascular grafts. Biomacromolecules, 2009. 10(4): 814±21. 158. Uttayarat, P., et al., Microtopography and flow modulate the direction of endothelial cell migration. Am J Physiol Heart Circ Physiol, 2008. 294(2): H1027±35. 159. Zhang, X., et al., Dynamic culture conditions to generate silk-based tissueengineered vascular grafts. Biomaterials, 2009. 30(19): 3213±23. 160. An, Y.H. and J.F. Richard, Concise review of mechanisms of bacterial adhesion to biomaterial surfaces. J Biomed Mat Res, 1998. 43(3): 338±48. 161. Katsikogianni, M., et al., Concise review of mechanisms of bacterial adhesion to biomaterials and of techniques used in estimating bacteria-material interactions. Eur Cells Mat, 2004. 8: 37±57. 162. Bollen, C.M., et al., The influence of abutment surface roughness on plaque accumulation and peri-implant mucositis. Clin Oral Implants Res, 1996. 7(3): 201±11.

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

163. McAllister, E.W., et al., The role of polymeric surface smoothness of biliary stents in bacterial adherence, biofilm deposition, and stent occlusion. Gastrointestinal Endoscopy, 1993. 39(3): 422±25. 164. Flint, S.H., J.D. Brooks, and P.J. Bremer, Properties of the stainless steel substrate, influencing the adhesion of thermo-resistant streptococci. J Food Eng, 2000. 43(4): 235±42. 165. Whitehead, K.A., J. Colligon, and J. Verran, Retention of microbial cells in substratum surface features of micrometer and sub-micrometer dimensions. Colloids Surfaces B: Biointerfaces, 2005. 41(2±3): 129±38. 166. Barnes, L.M., et al., Effect of milk proteins on adhesion of bacteria to stainless steel surfaces. Appl Environ Microbiol, 1999. 65(10): 4543±8. 167. Scardino, A.J., E. Harvey, and R. De Nys, Testing attachment point theory: diatom attachment on microtextured polyimide biomimics. Biofouling, 2006. 22(1): 55. 168. Hoipkemeier-Wilson, L., et al., Antifouling potential of lubricious, microengineered, PDMS ± elastomers against zoospores of the green fouling alga ulva (Enteromorpha). Biofouling, 2004. 20(1): 53. 169. Howell, D. and B. Behrends, A review of surface roughness in antifouling coatings illustrating the importance of cutoff length. Biofouling, 2006. 22(6): 401. 170. Verran, J., et al., Microbial retention on open food contact surfaces and implications for food contamination. Adv Appl Microbiol, 2008, 64: 223±46. 171. Emerson IV, R.J., et al., Microscale correlation between surface chemistry, texture, and the adhesive strength of Staphylococcus epidermidis. Langmuir, 2006. 22(26): 11311±21. 172. Cousins, B.G., et al., Effects of a nanoparticulate silica substrate on cell attachment of Candida albicans. J Appl Microbiol, 2007. 102(3): 757±65. 173. Rimondini, L., et al., The effect of surface roughness on early in vivo plaque colonization on titanium. J Periodontol, 1997. 68(6): 556±62. 174. Quirynen, M., et al., The influence of titanium abutment surface roughness on plaque accumulation and gingivitis: short-term observations. Int J Oral Maxillofacial Implants, 1996. 11(2): 169±78. 175. Tang, H., et al., Influence of silicone surface roughness and hydrophobicity on adhesion and colonization of Staphylococcus epidermidis. J Biomed Mat Res Part A, 2009. 88A(2): 454±63. 176. Wennerberg, A., et al., Some soft tissue characteristics at implant abutments with different surface topography. A study in humans. J Clin Periodontol, 2003. 30(1): 88±94. 177. Alava, J., et al., Effects of bacterial adhesion with respect to the type of material, structure and design of intraocular lenses. J Mat Sci: Mat Med, 2005. 16(4): 313± 317. 178. Tebbs, S.E., A. Sawyer, and T.S.J. Elliott, Influence of surface morphology on in vitro bacterial adherence to central venous catheters. Br J Anaesth, 1994. 72(5): 587±91. 179. Thomsen, P. and C. Gretzer, Macrophage interactions with modified material surfaces. Curr Opin Solid State Mat Sci, 5(2±3): 163±76. 180. Eriksson, C., J. Lausmaa, and H. Nygren, Interactions between human whole blood and modified TiO2-surfaces: influence of surface topography and oxide thickness on leukocyte adhesion and activation. Biomaterials, 2001. 22(14): 1987±96. 181. Murray, D.W., T. Rae, and N. Rushton, The influence of the surface energy and roughness of implants on bone resorption. J Bone Joint Surg ± Series B, 1989. 71(4): 632±7.

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182. Rich, A. and A.K. Harris, Anomalous preferences of cultured macrophages for hydrophobic and roughened substrata. J Cell Sci, 1981. 50(1): 1±7. 183. Tanaka, T., et al., Cell adhesion to acrylic intraocular lens associated with lens surface properties. J Cataract Refractive Sur, 2005. 31(8): 1648±51. 184. Refai, A.K., et al., Effect of titanium surface topography on macrophage activation and secretion of proinflammatory cytokines and chemokines. J Biomed Mat Res Part A, 2004. 70A(2): 194±205. 185. Brohim, R.M., et al., Early tissue reaction to textured breast implant surfaces. Ann Plastic Surgery, 1992. 28(4): 354±62. 186. Mohanty, M., et al., Evaluation of soft tissue response to a poly[urethane urea]. Biomaterials, 1992. 13(10): 651±6. 187. Parker, J.A.T.C., et al., Soft-tissue response to silicone and poly-L-lactic acid implants with a periodic or random surface micropattern. J Biomed Mat Res, 2002. 61(1): 91±8. 188. WaÈlivaara, B., et al., Titanium with different oxides: in vitro studies of protein adsorption and contact activation. Biomaterials, 1994. 15(10): 827±34. 189. Koh, L.B., I. Rodriguez, and S.S. Venkatraman, A novel nanostructured poly(lacticco-glycolic-acid)-multi-walled carbon nanotube composite for blood-contacting applications: thrombogenicity studies. Acta Biomaterialia, 5(9): 3411±22. 190. Hirata, I., et al., Deposition of complement protein C3b on mixed self-assembled monolayers carrying surface hydroxyl and methyl groups studied by surface plasmon resonance. J Biomed Mat Res Part A, 2003. 66A(3): 669±76. 191. Labarre, D., et al., Complement activation by substituted polyacrylamide hydrogels for embolisation and implantation. Biomaterials, 2002. 23(11): 2319±27. 192. Toda, M., et al., Complement activation on surfaces carrying amino groups. Biomaterials, 2008. 29(4): 407±17. 193. Bonfield, W., Hydroxyapatite-reinforced polyethylene as an analogous material for bone replacement. Ann New York Acad Sci, 1988. 523: 173±7. 194. Pla-Roca, M., et al., Micro/nanopatterning of proteins via contact printing using high aspect ratio PMMA stamps and nanoimprint apparatus. Langmuir, 2007. 23(16): 8614±18. 195. Birch, H.M., Cell culture in three dimensions. Nature, 2007. 446: 937. 196. Engler, A.J., et al., Matrix elasticity directs stem cell lineage specification. Cell, 2006. 126(4): 677±89.

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Techniques for analysing biomaterial surface chemistry J . Y A N G and M . R . A L E X A N D E R , The University of Nottingham, UK

Abstract: This chapter presents an introduction to the analytical methods that are most commonly used to chemically analyse the surface of materials. The content and level of detail is aimed at the non-expert as an accessible introductory guide to the principles and practical application of surface analysis in the field of biomaterials modification. Key words: surface chemical analysis, x-ray photo electron spectroscopy (XPS), time of flight secondary ion mass spectrometry (ToF SIMS).

8.1

Introduction

In the majority of cases the surface chemistry sensed by `the biology' with which biomaterials are designed to interact will be in the uppermost few nanometres of the material. Consequently techniques that provide information on the chemical structure of surfaces from within that depth are required. The routine method for information from surfaces in Materials Science is electron microscopy, which coupled with energy dispersive X-ray analysis provides chemical and topography information typically from the top few microns. This is 1000 times deeper than is of relevance to biological surface interactions. Fortunately surface chemical techniques exist that routinely obtain chemical information from the uppermost 1±10 nm of surfaces, these being X-ray photoelectron spectroscopy (XPS) and secondary ion mass spectrometry (SIMS). XPS was developed in the 1950s by Kai Siegbahn and SIMS in the late 1960s by Alfred Benninghoven, and both were rapidly applied by the semiconductor industry in the drive for electronic miniaturisation which still exists today. The application of these techniques to biomaterials was pioneered by Buddy Ratner and and co-workers at the University of Washington.1±5 Characterisation of the surface chemistry is necessary if the properties and performance of the native or modified surface of any material is to be understood. The surface elemental or functional composition can be quantified using XPS, while structurally rich but non-quantitative mass spectral information can be obtained from SIMS. When such tools are not readily available, it is often

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assumed that the starting material is as-described by the supplier and the reaction of the proposed surface modification scheme holds. Since both of these information sources normally originate from solution or solid bulk analytical characterisation, this is invariably a mistake. Surfaces are very prone to contamination; extraneous compounds may arrive at the surface through contact with liquids or other solids, adsorption of volatile species from the atmosphere, or segregation of contaminant from the bulk to the surface.* Amounts of contaminants that are insignificant when measured as a proportion of the bulk of the material can dominate when concentrated at a surface. Surface chemistry laboratories that analyse samples from a range of sources will attest to the fact that a very large proportion of their analyses include contaminants. In contract to surface analysis laboratories, this is often the reason for the failure of the component or process in question and remedial action is taken as a result of the finding. However, when the goal is the production of chemically well-defined surfaces for the development of devices that will have applications in healthcare, it is an unnecessary complication that must be resolved by the use of appropriately sourced starting materials and suitable preparation methodologies. This is often a trial and error process which can require surface analysis at various points in multi-step surface modification processes.6 Furthermore, the batch to batch reproducibility and the spatial homogeneity of the surface chemistry of material after modification needs to be confirmed since unexpected inter and intra sample variance often arises. Surface analysis facilities require significant capital and continued infrastructure investment, this requirement can be daunting for a material chemist. However, since surface analysis equipment is ideally operated by a skilled expert and maintenance costs are high, those who run it in academia normally provide access to their facilities, often with research council funding for academics.y There are also a number of commercial surface analysis firms who specialise in selling expertise and equipment access.z It is worth briefly mentioning water contact angle (WCA) measurement, since although not strictly chemical analysis, it is probably the oldest method used to follow changes in surface chemistry through the surface property of wettability.7 This is the most rapid and cost effective means of detecting changes in surface chemistry from the top nanometre of solid surfaces. It relies on detecting changes in the surface energy influencing the contact equilibrium with water; in the most common measurement, the sessile drop experiment, a water drop is placed upon the surface and its angle of contact with the surface at the * y z

Best practice sample handling procedures are described in Section 8.4, that minimise the chance of contamination during preparation, storing and transportation. www.uksaf.org/home.html, http://www.nottingham.ac.uk/nano/OpenAccess/openaccess.phtml, http://www.nb.engr.washington.edu/ http://www.csma.ltd.uk/, http://www.eaglabs.com/, http://www.uwo.ca/ssw/, http:// www.oxfordsurfaces.com/, http://www.molprofiles.co.uk/

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perimeter of the drop is estimated. Whilst this is very surface sensitive, it does not provide any specific information on what causes any change, and consequently it is often used as a screen to get a quick estimate of the efficacy of a modification without the need to enter the sample into a vacuum spectrometer in combination with a more detailed XPS and SIMS analysis. When the identity of the sample is well understood in a model system, WCA can give additional information unavailable by XPS and SIMS, e.g. the molecular order of self assembled monolayers.8±11 When multiple liquids are employed an estimate of the surface energy can be obtained.12

8.2

X-ray photoelectron spectroscopy (XPS)

XPS is the most commonly applied surface chemical analysis technique*, which has been used to characterise elemental composition and chemical functional groups of materials increasingly intended for biomedical application. XPS, also referring to ESCA (electron spectroscopy for chemical analysis), allows quantitative elemental analysis, except for hydrogen and helium, of the uppermost surface (approximately 10 nm) of materials. Chemical state information is provided by utilising the chemical shift effect. These two important quantitative insights into the surface chemistry are the reason that XPS has become so widely applied.

8.2.1

The electron photoemission process

XPS relies on the photoelectric effect, which was initially described by Einstein for which he received the Nobel Prize in 1921. When a beam of photons irradiate a solid surface some interact with an atomic orbital electron with total transfer of the photon energy to the electron. If this energy is greater than that required to free the electron from the hold of the nucleus, then a photoelectron is emitted bearing any excess energy imparted by the photon as represented in Fig. 8.1. An energy balance relates the amount of excess kinetic energy (Ek) measured for the photoelectron, to the binding energy (Eb) of the electron to the nucleus, the work function () and the input photon energy h (h is Planck's constant and  is the frequency of the X-rays employed): h ˆ Eb ‡ Ek ‡  The work function, , is the energy required to free the electron from the surface to the vacuum and is specific to the spectrometer and material under analysis.  is calculated when calibrating the energy scale of the spectrometer. Thus, for an X-ray source of fixed energy, measurement of Ek of the emitted photoelectrons *

Search topic `XPS or ESCA' in ISI Web of Knowledge results for all years on 20/8/9 returns 63,259 results and `ToF SIMS or static SIMS' returns 3,334.

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8.1 Photoelectron and Auger electron emission processes.

readily allows Eb to be calculated from which identification of the atomic orbital can be made. All energies are expressed in electron volts (eV). The vacancy left by the ejected photoelectron can be filled by an outer orbital electron from a higher energy level and the excess energy gained by the atom from this transition is transferred to another outer orbital electron which results in its emission as show in Fig. 8.1. This process is called the Auger process and these Auger electrons are observed in XPS spectra as broad peaks on a background step. Auger electrons stimulated by electron bombardment are utilised in Auger Electron Spectroscopy (AES), a related surface analysis technique which is restricted to the analysis of conducting surfaces. In XPS, the atomic orbital core levels from which the photoelectrons are ejected are described using the quantum numbers for the core level (Table 8.1). Different atoms emit photoelectrons at characteristic binding energies, which enables identification of the elements. The photoelectron intensity is proportional to the amount of the corresponding element; therefore, by measuring the area of the peak the elemental composition of the surface can be calculated. XPS is a surface sensitive technique because photoelectrons can only travel a short distance in solids without losing energy in an inelastic collision with a surface atom, although the soft X-ray photons commonly used actually penetrate many microns into the sample. Table 8.1 Atomic orbital nomenclature Principle quantum number (n) Angular momentum quantum number (`):

n  1, 1, 2, 3, 4 . . . or K, L, M, N . . . 0  `  n ÿ 1, 0, 1, 2, 3 or s, p, d, f

Spin quantum number (ms)

Au4f5/2 Au4f7/2

ÿ! ÿ!

ÿ! Element

n `

ÿ!

Example:

ÿ1/2 or 1/2

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Spin-orbit splitting: ` ÿ 1=2 or ` ‡ 1=2

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Instrumentation The main components of an XPS spectrometer are illustrated in Fig. 8.2. The experiment is carried out under ultra high vacuum to prevent surface contamination by atmospheric species, to enable high voltages to be utilised and to prevent photoelectrons undergoing inelastic collisions in the spectrometer prior to detection. To generate X-rays in a laboratory instrument, a heated filament is used to produce electrons which are then accelerated to a target by a potential, typically 10±20 kV. Bombardment of the X-ray anode by the high-energy electrons results in the emission of X-rays. The X-ray anode is usually aluminium or magnesium producing Al K X-rays with energy h ˆ 1486:6 eV or Mg K Xrays h ˆ 1253:6 eV. Most modern XPS spectrometers are equipped with aluminium anodes with a monochromator to diffract the 1486.6 eV X-ray line, thereby narrowing the energy spread, excluding emission at other energies and allowing the illumination area to be defined. Monochromatic X-rays generate spectra with lower background and narrower peaks compared to earlier nonmonochromatic instruments that were usually equipped with a dual Mg/Al anode. A lens system is used to collect photoelectrons and pass them to an electron energy analyser that discriminates electron energy by scanning the voltage between two hemispherical surfaces to provide a serial analysis of the

8.2 Schematic of an XPS instrument.

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electrons from the sample. Traditionally, electron detection was carried out by photomultipliers, but recent delay-line detectors have been developed for commercial XPS instruments with the aim of providing quantitative imaging.13

8.2.2

Elemental quantification

The area of a core level representing each of the elements in an analysis is used to calculate the elemental composition of the analysed area assuming a lateral and vertically homogeneous distribution.* To measure the area of a photoelectron peak, a background is needed to separate the electrons that have undergone inelastic scattering from those that have not. A linear background is normally used for polymers, although for other materials a large intensity increase is often observed over the range of the peak, necessitating the use of other backgrounds such as Shirley or Tougaard.14 To achieve accurate atomic concentrations, the measured peak areas need to be scaled differently to account for the differences in probabilities of photoelectron creation and collection efficiency by the spectrometer. The basic equation for calculating an elemental concentration is: Ni;k ˆ I0 i i;k i;k Ti;k where Ni,k is peak area determined by experiments for the kth orbital of element i; I0 is the X-ray flux. i is the volume density of element i in the total volume examined by XPS, which consequently derives the atomic concentration of element i in the examined surface volume. i;k is the mean free path. i;k is the photoionisation cross section which is the probability of the incident X-ray creating a photoelectron from the kth orbital of element i. The photoionisation cross section for the particular element and orbital has been calculated and tabulated by Scofield.15 The Scofield cross sections are used to theoretically calculate the relative sensitivity factors (RSFs) in an approach offered by the National Physical Laboratory (NPL) UK to quantify atomic concentrations.16 The transmission function, Ti,k, of the instrument expresses the collection efficiency of ejected photoelectrons. The transmission function may be calculated by measuring the spectrum of copper, silver and gold standards and referenced to the NPL metrology spectrometer.17y Data is normally acquired using a survey scan conducted under maximum spectral sensitivity acquisition conditions to find all elements present on the surface over the full range for the X-ray source in use (Fig. 8.3). When more than one core level is detected for each element, the most intense peak is usually * y

For guidance on procedures for when the surface is vertically heterogeneous, which is the most common situation when surface modification is concerned, see section entitled Depth distribution. http://www.npl.co.uk/nanoscience/surface-nanoanalysis/products-and-services/referencematerials

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8.3 XPS survey scan of poly(L-lactic acid).

chosen to measure the composition to maximise the precision. High energy resolution scans are acquired from core levels of interest under acquisition conditions chosen to obtain optimal energy resolution.

8.2.3

Chemical state information

The chemical shift effect results from the influence of the chemical environment on the binding energy of the detected photoelectrons. This phenomenon is useful for distinguishing different chemical functionalities on surfaces as can be seen for the C1s core level for different organic functionalities in Table 8.2.18 Highresolution spectra are used to identify and accurately measure the peak areas of different chemical states. XPS spectra may be processed using the software provided by the instrument supplier or a generic processing tool.* Curve fitting of overlapping peaks from different chemical environments that contribute to core levels is the most challenging aspect for those new to XPS *

CasaXPSÕ is a very popular instrument independent XPS processing package.

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Table 8.2 Typical C1s binding energy for organic samples. Reprinted with permission from reference 18. Copyright Wiley (2009) Functional group

Hydrocarbon Amine Alcohol, ether Cl bound to carbon F bound to carbon Carbonyl Amide Acid, ester Urea Carbamate Carbonate 2F bound to carbon Carbon in PTFE 3F bound to carbon

C±H,C±C C±N C±O±H, C±O±C C±Cl C±F C=O N±C=O O±C=O N±C(=O)±N O±C(=O)±N O±C(=O)±O ±CH2CF2± ±CF2CF2± ±CF3

Binding energya (eV)

Chemical shift (eV)

285.0 286.0 286.5 286.5 287.8 288.0 288.2 289.0 289.0 289.6 290.3 290.6 292.0 293±294

± 1.0 1.5 1.5 2.8 3.0 3.2 4.0 4.0 4.6 5.3 5.6 7.0 8.0±9.0

a

The observed binding energies will depend on the specific environment where the functional groups are located. Most values are quoted to a precision of 0.2 eV, but some (e.g., fluorocarbon samples) can be larger.

analysis. This process involves fitting a number of synthetic components to represent the contribution of individual chemically shifted components to the peak envelop. For organic samples, the core levels containing the most useful information are the C1s and the O1s. After insertion of a background, the synthetic peaks are added to represent the individual chemical environments. Their shape is usually well described by summed Gaussian±Lorentzian (GL) functions, although it is sometimes necessary to introduce high binding energy asymmetry into peak shapes to account for processes such as vibrational fine structure (e.g., resolved in poly(ethylene).19 In Fig. 8.4 the example of PLA is presented in which three chemical states of C1s can be easily distinguished. These three chemical states of C1s in PLLA are hydrocarbon, ether and carboxyl as labelled in the figure. The three peaks of C1s were fitted using the GL function, respectively. The parameters used for the fitting are shown in Fig. 8.4.20 The goodness of fit may be assessed visually or quantitatively using a single parameter such as chi-square goodness-of-fit or a plot of the residual versus binding energy to highlight areas of poor fit. Normally the addition of more peaks improves the fit to which the software iterates, but it does not necessarily improve the accuracy of the surface quantification. It is therefore good practice to have a specific reason for peak introductions such as the different chemical environments in the structure of the PLA (Fig. 8.4). The full width at half maximum (FWHM) is a commonly used measure of peak width. Within one core level, the FWHM of peaks from the same phase are generally

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8.4 C1s core level of PLA. Adapted with permission from Reference 20. Copyright Elsevier (2009).

maintained to be about equal. In this example the component model peaks are well defined by the spectrum and their position and relative intensities are therefore easily measured, although this is not always the case. Assignment of core level components is greatly aided by reference to spectra from standard polymers.19 Certain functional groups are associated with the same chemical shift, which makes the identification and quantification of the functional group difficult. For example, carbons bound in an ether environment (C±O±C) and carbons in a hydroxyl environment (C±OH) are both observed at approximately 286.5 eV. Also, carbons in a carboxylic acid, C(=O)OH, environment and carbons in an ester environment, C(=O)OR, are not readily distinguishable based upon binding energy shifts. In this case, a priori knowledge on the structure of the analysed

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sample is very helpful to assign chemical states correctly because it guides the components to include in the peak envelope. If this information is not available, other techniques such as IR are needed to assist the identification of specific functional groups.21 A chemical reaction specific to only the functional group of interest can be carried out to assist identification of specific functionalities in a process called chemical derivatisation.22,23 This reaction should identify the functional group by altering its chemical shift or by adding a tag atom not present on the surface prior to reaction.

8.2.4

Depth information in XPS

Inelastic mean free paths (IMFP) Photoelectrons do not travel far in solids due to interactions with atoms resulting in inelastic scattering. The electrons that undergo energy loss through inelastic collisions are observed in the spectra as steps increasing in intensity to the high binding energy side of the core levels (Fig. 8.3). The number of photoelectrons that reach a surface without suffering energy loss through collisions are described by the Beer-Lambert law: N ˆ N0 eÿt= sin 

8:1

where  is the IMFP which is the average distance that an electron with a given energy travels between successive inelastic collisions, N0 and N are the number of electrons before and after passage through the thickness of sample, t, while  is the angle between the surface and the plane of acceptance of the electron optics. When t ˆ , N ˆ N0 eV; i.e. 63% of the photoelectron have lost energy and no longer contribute to the peak in XPS spectrum. The depth from which the XPS signal is derived is normally defined as the average depth normal to the surface from which 95% (t ˆ 3 in Equation 8.1) of the ejected photoelectrons are derived. The inelastic mean free path is a function of kinetic energy, i.e. depending on the X-ray source. Clark et al. first described the relationship between mean free path and kinetic energy of the photoelectron to be  / Ek0:5 using a wide kinetic energy range (500±4500 eV).24 A database of IMFPs in solid elements and compounds can be found in literature.25 The mean free paths can be determined experimentally using films with known thickness to detect the attenuation of the core level from a specific element only existing in the substrate.26±29 Equation 8.1 can be used with literature values of  to measure the thickness of an overlayer by measuring the attenuation of the signal from the substrate. However, using effective attenuation length (EAL) rather than IMFP () has been proven to generate more accurate thickness measurements due to the consideration of elastic electron scattering.30,31 In order to carry out the overlayer thickness measurement, the substrate has to contain a marker element or chemical environment that is not present in

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the coating provided the coating is thinner than the analysis depth of the marker element. This technique has been widely used with various forms of Equation 8.1, e.g. to measure a plasma polymerised hexane layer with varying thickness on plasma polymerised allylamine (ppAAm) coated poly(methyl methacrylate) substrate.32 The N1s core level in the ppAAm was used as the marker. The peak area of N1s from uncoated ppAAm was first measured and used as the N0 value in Equation 8.1. Then peak areas of N1s from coated ppAAm by ppHex with varying thicknesses were measured, and these values are N in Equation 8.1. For the mean free path of N1s as 2.8 nm, determined from the literature,33,34 the effective thickness of the ppHex layer was measured to be between less than 1 nm and 8 nm across 10 mm of the thickness gradient. In the absence of a marker element in the substrate, for instance, to measure SiO2 thickness on Si, a modified equation must be used including the mean free path of Si 2p in the oxide and the ratio of the intensities of the Si 2p peak from bulk SiO2 and from pure Si.35 The equation for calculating the thickness of the oxide layer is: doxide ˆ LSiO2 …ESi †cos  ln…1 ‡ Rexpt =R0 † where LSiO2(ESi) is the attenuation length for the Si 2p photoelectrons in SiO2,  is the angle of emission of the detected electron from the surface normal. Rexpt is 1 1 1 1 =ISimet . ISiO and ISimet the measured ratio ISiO2 =ISimet and R0 is the ratio of ISiO 2 2 are the Si 2p intensities of corresponding pure bulk materials. ISiO2 and ISimet are the intensities of the Si 2p peak in the oxide and metallic states. Depth information The distribution within the XPS analysis depth is of great interest in studies involving surface modifications.36±40 A profile of the chemical distribution versus depth can be achieved by destructive or non-destructive means. A common destructive method is surface erosion by ion sputtering in which the surface is bombarded with energetic ions which causes removal of the surface atoms by sputtering. The newly exposed surface can then be analysed. Several sputtering techniques for depth profiling have been developing to meet the requirements of different classes of materials. Low degradation sputtering using buckminsterfullerene and other polyatomic sources have recently become available to analyse organic multilayer structures in which the chemical composition was found to be preserved.41 Extremely low sputtering degradation of polytetrafluoroethylene was also observed when using C60 ion beam to remove a deliberately introduced contamination layer on the polytetrafluoroethylene surface.42 The sources have recently become available commercially on both XPS and time of flight secondary ion mass spectrometry (ToF SIMS) instruments. Non-destructive depth profiling can be achieved by angle resolved XPS, a method which uses the surface sensitivity of XPS. By tilting the sample, i.e.

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8.5 Depth profiling by reducing the take-off angle.

reducing the take-off angle (), the photoelectrons collected in the plane of the analyser have to travel a greater distance within the solid before emerging from the sample (Fig. 8.5). Thus, the depth from which photoelectrons are detected without undergoing inelastic collision decreases with ;  ( ˆ 90) > t1 …1 † > t2 …2 †. There are various algorithms that aim to reconstruct a depth profile from angle resolve XPS data which have been reviewed by Cumpson, many of which are available in a software package called ARCtick from NPL.43,44 The shape of the background on the high binding energy side of a core level contains information on the depth of the core level atoms. If the atoms are buried below other material, more photoelectrons will be inelastically scattered than if at the surface. There will, therefore, be a rising background composed of the photoelectrons that have lost energy on their way to the surface, and no longer appear in the photoelectron peak. In many analyses, this qualitative observation is used to define the depth of the component satisfactorily. However, quantitative analysis of the background shapes have been made possible by detailed measurements on electron losses in materials which are employed by Tougaard in software called QUASES.45,46 This approach enables quantitative overlayer measurements to be made and information on surface structures to be determined from survey spectra acquired at only one take off angle.

8.2.5

XPS imaging

XPS imaging may be achieved either with a small spot X-ray source by rastering the beam position, or using a position sensitive electron detector. Since X-rays are difficult to focus and detectors have become relatively advanced, the spatial resolution of scanning XPS is generally poorer than when using a position sensitive detector.47±49 Spatial resolution of a few microns can be achieved using this method. An example of an XPS image is shown in Fig. 8.6,50 where the presence of patterns of plasma polymerised allylamine was demonstrated by imaging of the nitrogen (N 1s) distribution using parallel photoelectron imaging. Quantitative XPS images cannot reliably be produced by counting at single binding energies since small shifts due to lateral differential charging can greatly

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8.6 The XPS imaging (N 1s) of strips of plasma polymerised allylamine on a plasma polymerised 1,7-octadiene coated silica substrate. Reprinted with permission from Reference 50. Copyright Wiley (2001).

change peak intensity estimates. Instead, acquisition of intensity maps at small energy steps over the core levels of interest are necessary to construct spectra from the images.51

8.2.6

Other spectral features

Surface charge Due to the emission of photoelectrons, the sample will charge positively if it is an insulator, whereas conductive samples gain electrons from the spectrometer through contact with the sample holder and develop no charge. Photoelectrons must overcome the potential barrier produced by the positive charged surface to emit into the vacuum and be transferred on to the analyser. This makes the corresponding peaks appear at a higher binding energy. Moreover, the charging is usually not uniform for a variety of reasons including non-uniform X-ray illumination and sample topography. A surface with non-uniform charge will provide broadened peaks. In order to compensate for the positive charge, XPS instruments use an electron source to compensate for the charge built up on the insulator surface with a stream of low-energy electrons. Since the `charge correction' in the spectrometer is never exact, the spectra from insulators are charge referenced to a defined point, which for polymers is normally chosen to be the hydrocarbon environment in the C1s at 285.0 eV. Using this reference point the binding energy axis is normally adjusted prior to data processing.

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Shake-up satellites Satellite peaks on the high binding energy side of the core level arise from a specific amount of energy loss of the outgoing photoelectrons. Such phenomena include shake-up and shake-off. The shake-up process occurs when a valence electron is excited to a higher energy level by energy transferred from the ejected photoelectron. If sufficient energy is transferred to a valence electron, the electron can finally leave the atom completely; this process is called shakeoff. The C1s chemical shift between alkyl and unsaturated or aromatic bonds is small and difficult to detect by XPS under normal conditions. The shake-up satellite can be used as a diagnostic for the presence of unsaturated bonds at the polymer surface. For example, the shake-up process has been observed in XPS spectrum of poly(benzylmetharylate) which is due to - molecular orbital transitions (Fig. 8.752). The shake-up processes in polymers were studied extensively by Clark, Dilks, and coworkers.53±57 Doublets Certain core levels yield two peaks, e.g. the Au4f splits into two sub-peaks Au 4f5/2 (Eb ˆ 87.5 eV) and Au 4f7/2 (Eb ˆ 83.8 eV) (Fig. 8.818); this effect is observed due to a magnetic interaction between the spin of the electron (up and down) and its orbital angular momentum, which leads to a splitting of the energy

8.7  bond shake up of C1s from XPS of poly(benzylmetharylate) containing aromatic and other unsaturated groups, showing the presence of a discrete shake-up satellite at 6.3 eV from the main peak. Reprinted with permission from Reference 52. Copyright Springer (1985).

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8.8 Spin-orbital coupling leads to splitting of 4f photoemission of gold into two sub-peaks. Reprinted with permission from Reference 18. Copyright Wiley (2009).

level to two components. This is particularly pronounced for transition metals, such as gold and iron. This interaction is referred to as spin-orbit coupling. The ratios of intensities of common spin-orbit pairs are 1 : 2 (p1/2 : p3/2), 2:3 (d3/2 : d5/2) and 3:4 (f5/2 : f7/2).

8.3

Time of flight secondary ion mass spectrometry (ToF SIMS)

SIMS was first developed for the analysis of inorganic materials such as semiconductors, but relatively quickly the potential in organic systems was realised and it was applied to polymers and organic small molecules such as amino acids, peptides and drugs.58,59 Importantly, these early studies on organics established a link between the molecular structure and the mass spectrum generated by SIMS, which to many was unexpected because the process was based on sample damage (sputtering) and of the association of SIMS with inorganic analysis. This misapprehension has largely been dispelled as the vast literature demonstrating the application of SIMS to provide valuable surface information continues to grow.

8.3.1

Principles

SIMS is based on the mass analysis of charged atoms and molecules generated by the impact of energetic primary ions with the surface (Fig. 8.9). The primary

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8.9 Schematic of collision between primary ions and a surface.

ions generate a collision cascade that involves energy transfer from the impacting primary ions to the material. The species sputtered from the surface in this process include neutral, positively and negatively charged atoms and molecules. Only a small portion (10±7 to 10±3)60 of these species are charged and can therefore be directly detected. Methods like laser post-ionisation secondary neutral mass spectrometry have been successfully developed to enhance the useful yield (atoms detected per atom sputtered) by ionising neutral fragments, although they are not widely applied. Newer primary ion cluster beams, such as gold, C60 and bismuth have most recently been applied to improving yields over earlier sources such as caesium and gallium. In order that the ions analysed are representative of the undamaged surface, a maximum primary ion dose has been defined, which is 1013 ions/cm2 called the static limit, below which the probability of a secondary ion resulting from previously sputtered areas is low.61 Time of flight secondary ion mass spectrometry (ToF SIMS) is the most common type of instrument for SIMS analysis within the static regime. A pulse of primary ions is used to produce a packet of secondary ions. To achieve high mass resolution, pulsed primary ions are bunched to reduce the pulse duration by compressing the primary ions in space. Ions of different masses are differentiated by the different flight time over a set distance when accelerated to a constant potential. Ions are first accelerated to the same kinetic energy via an electric field (Ua) before entering the field-free drift region. Before reaching the detector, they travel the same distance, thus the speed of ions travelling in the drift region depends on their masses according to the equation: Ekin ˆ m 2 =2 ˆ eUa

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8.10 Schematic of the energy compensation in a single stage mirror reflectron. The relative position of secondary ions of high (ú) and low energy (n) from the sample to the detector is indicated at constant time intervals. The dashed line indicates the trajectory of the ions. The ions are assumed to start from the same spot at the same time normal to the sample surface.

As the distance of the drift region is fixed, the time spent by a specific ion reaching the detector can be converted into its mass. In practice, a small range of initial kinetic energy of secondary ions exists. To combat this, a reflectron analyser is adopted to increase the flight path of high-energy ions with respect to the lower energy ions. A schematic of a single stage reflectron is shown in Fig. 8.10. The system consists of a field free drift region of length L1, a homogeneous retarding mirror field of length Dm and a second drift region of length L2. To reflect the secondary ions to the detector, the retarding potential of the mirror has to satisfy eUm  eU0 , where eU0 is the nominal secondary ion energy. As the high-energy ions travel a longer distance with respect to the lower energy ions, the flight time dispersion due to the initial kinetic energy dispersion is compensated. For insulating materials such as polymers, surface charging results due to charge transfer from primary ions, generation of secondary ions and electrons. Without correction this would significantly deteriorate the quality of the secondary ion mass spectra. An electron flood gun is applied to bathe the surface in low energy electrons to correct charge imbalances at the sample surface. These electrons are emitted during the interval of extraction potential and are optimised by the operator. By application of a field of appropriate polarity, ions can be extracted and ToF mass analysed to produce both a positive and a negative ion mass spectrum. A positive ion spectrum of polyglycolic acid (PGA) is shown in Fig. 8.11.62 In this positive ion spectrum, ions containing up to 4 monomer units (M) can be seen.

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8.11 A positive ion static SIMS spectrum from PLA acquired using Cs+ primary ions. Key: `' ˆ [nM+H]+ (n ˆ 2±4), `·' ˆ [nM-OH]+ (n ˆ 1±4) and `y' ˆ [2M± CO2±H]+. Reprinted with permission from Reference 62. Copyright Elsevier (2006).

8.3.2

Spectrum interpretation

As a result of surface entanglement, SIMS of polymers routinely produces numerous fragments that rarely include a molecular ion. This can be seen in Fig. 8.11 where the largest detected fragment is only 4 monomer units in size. Small molecules, such as drugs or surface immobilised compounds, often produce a molecular ion which can readily be assigned based on comparison with knowledge of the molecular identity of the surface. When the surface components are not known, assignment of peaks in SIMS spectra can be challenging and is certainly not as routine as conventional mass spectra due to the high degree of fragmentation often encountered, producing hundreds of peaks from a single chemical entity. Although for convenience hydrogen is widely regarded as having a mass of 1 u (unified atomic mass unit), in fact its exact mass is 1.00783 u, and most other elements have a non-integer mass when considered to three decimal places. Modern instruments can accurately resolve such differences and measure peak positions to this accuracy allowing stochimetric assignments to be made using the accurate mass (Fig. 8.12). Despite this capability, assignments of single peaks are often inconclusive and complementary assignments from the same or other ion spectrum are useful. This is aided by comparison of spectra with

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8.12 High mass resolution spectrum of PLA around m/z 71.

spectral libraries. Handbooks of mass spectra of compounds have been published from the early days of SIMS, which are still being joined by new works including more compounds that are used to aid the analyst in identification of compounds from unknown samples.63 The SurfaceSpectraÕ version 4 has included over 1900 spectra and more than 1000 materials.* Most of the spectra included were acquired from high resolution ToF instruments. Some general guidance of spectral interpretation for organic materials follows. Materials where molecular species are easily desorbed from the surface and have a readily ionisable group can exhibit an ion close to the mass of the intact molecule. Materials containing hydrocarbon structures, the simplest example being polyethylene, exhibit a fingerprint spectrum of peak clusters with each ion labelled CnHm‡. The characteristic peaks in negative spectra from nitrogen containing materials are CNÿ and CNOÿ (if combined with oxygen). Fragments containing one nitrogen atom are also notable by their tendency to appear at even masses. Aromatic hydrocarbons produce a fingerprint set of fragments including a prominent peak at m/z = 91 (C7H7‡, the tropillium ion). Unless they are very low molecular weight or contain an oligomeric fraction, polymers normally only exhibit the monomer unit, sometimes a dimer or slightly greater. Polyesters, which contain oxygen in the backbone such as poly(lactone), produce series such as [nM+H]‡, [nM-OH]ÿ, and [nM+OH]ÿ, with R±C=O‡ or R±COOÿ structures.64 Vinyl polymers containing oxygen in the side chain produce dominant ions from the side chain fragments, especially R±C=O‡.65 Acrylate polymers show characteristic positive ion at m/z 55(C3H3O‡) while methacrylates generate a characteristic ion with mass of 69 (C4H5O‡). Halogen containing polymers are characterised by intense negative ions from the halogen atom. *

http://www.surfacespectra.com/

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The peak intensity of ions of interest can be used to provide a qualitative estimate of the relative amount of material. However, since the ionisation probability and, therefore, the ion intensity may be heavily influenced by surrounding elements (the matrix effect), SIMS is not inherently quantitative. Furthermore, the number of ions selected for such comparison is small relative to the number of ions that originate from each surface component and are often arbitrary. This means that the interpretation can be influenced by the bias of the analyst and unanticipated features can be easily missed. To overcome the problem of the large amount of peaks requiring consideration, a multivariate approach by using principal component analysis (PCA) is useful to enable all peaks in the spectra to be considered. PCA can reduce a large data set, such as SIMS spectra in which each ion is considered as a dimension, to the smallest number of dimensions expressing the main variation in the dataset with minimal loss of information. One can think of this approach as reducing the hundreds of ions usually presented in SIMS spectra to a few number of `synthetic ions' which are called the principle components (PCs). Several software packages contain a PCA function, e.g. EigenvalueÕ.* A more detailed introduction to the mathematical background of PCA and another multivariate technique called partial least squares (PLS) regression can be found elsewhere.66,67 A good example of the use of PCA is presented in Fig. 8.1368 which shows the comparison of SIMS spectra of different proteins adsorbed onto a mica surface using PCA. It is apparent that the PCA approach can be used to distinguish between the different proteins adsorbed onto mica using two PCs representing the SIMS data. To provide a quantitative analytical capacity using SIMS data, e.g. correlating surface chemistry presented in SIMS spectra to cellular attachment, PLS regression can be used to build a model to obtain correlation between the ion intensities and the biological response. Chilkoti et al. have used PLS to predict bovine arterial endothelial cell attachment to 15 different plasma deposited organic films69 where good agreement between the predicted and experimental data was achieved (Fig. 8.1469). In this case the PLS regression was used to identify the ions controlling the cell growth and therefore identify surface moieties from which mechanistic understanding of the process can be built. Urquhart et al. have used PLS to predict water contact angle (WCA) from a library of 576 polymers in a microarray format.70 Good agreement between the prediction based on the PLS regression vector built from the SIMS data and experimental of the WCA was found (Plate III between pages 208 and 20970). Ions including oxygen and nitrogen were found to be associated with reduced water contact angle and hydrocarbon ions were associated with increased water contact angle. This makes physicochemical sense and, therefore, provides

*

http://www.eigenvector.com/

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8.13 Score plot from PCA of positive spectra of proteins adsorbed onto mica from 100 g/ml single component protein solutions. Reprinted with permission from Reference 68. Copyright American Chemical Society (2001).

8.14 Predicted BAEC growth by PLS versus measured values. Reprinted with permission from Reference 69. Copyright American Chemical Society (1995).

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confidence in the PLS methodology applied to the micro array sample format with a large number of samples. This multivariate methodology is finding greater application and is predicted to become routine for analysis of SIMS data.

8.3.3

ToF SIMS image analysis

The charged primary ion beam is readily focussed to a sub micron spot for most primary ion sources. ToF SIMS has therefore been widely used to provide surface mass spectral images. In more recent instruments, whole spectra can be rapidly collected and stored for each pixel, termed hyper spectral imaging. This capability allows the sample to be analysed without making prior assumptions to record certain spectral features of interest. To identify the distribution of different compounds, maps of characteristic ions fragmented from the compound of interest can be formed. In Fig. 8.15,71 ToF SIMS was used to image patterned alkanethiols on gold. The pattern was created by micro-contact printing using a PDMS stamp covered with tetraethylene glycol thiol (HS±(CH2 ) 11 ± (OCH2CH2)4±OH). Problems can arises due to the arbitrary selection of peaks used to construct SIMS image from the hundreds available since this process can be biased by the analyst. One approach to avoid this is to use multivariate technique such as multivariate curve resolution (MCR) to consider all the peaks present in the spectra.72 A good introduction to this method can be found in reference.18

8.15 ToF SIMS image of patterned tetraethylene glycol thiol (HS±(CH2)11± (OCH2CH2)4±OH) SAM on gold. The image shows the Oÿ ion distribution, which represents the distribution of the tetraethylene glycol thiol. Reprinted with permission from Reference 71. Copyright Elsevier (2003).

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Sample preparation and handling

An ISO standard is available for guidance on the handling of specimens and the containers for transport of sample intended for surface chemical analysis.73 In brief, it is best to avoid any contact with the sample surface that is to be analysed. If contact is unavoidable, it should be with a clean surface, and it is important to analyse the inner surface of the container by ToF SIMS to detect the presence of contamination. One commonly used material to store ToF SIMS samples is aluminium foil due to its ready availability, flexibility as a wrapping material and low levels of transferable contamination. Common industrial contaminants are silicones or plasticisers such as Di-iso octyl phthalate (DOP). Neither of these materials were found by ToF SIMS on foil produced commercially for domestic use.74 New glass vials rinsed carefully with high purity solvent can also be used to store samples. Care must be taken with the cap; polyethylene is ideal, a waxed card insert is undesirable. When such vessels are used to hold liquids in which surface modification is carried out, more stringent cleaning such as piranha etching if often employed.75 Common surface contaminants are mould-release agent used in the manufacture of injection-moulded plastic products such as poly(dimethyl siloxane) (PDMS) and plasticisers used in polymers such as phthalates. These are very evident in SIMS due to their strong molecular signals, but in XPS will only be apparent if at a high enough level to be detected through a unique elemental maker such as silicon. Storage products such as self-sealing bags contain such contaminants which can migrate over the sample surface with time and should be avoided. Characteristic peaks in ToF SIMS mass spectra of PDMS are at mass 28, 43, 73, 133 and 147. For DOP characteristic ions are 57, 71 and 149.63,76 It is of importance for the analyst to be aware of such characteristic peaks for surface contamination. Double-sided tape is often used to secure samples to the sample containers. Most are coated with PDMS since it is used to encourage release from the roll. Over time the PDMS can diffuse onto the sample surface, therefore, quick analysis after mounting or avoiding the use of PDMS tape is recommended. Samples in powder form can be prepared by adhering them to non silicone double-sided tape. This is achieved by allowing the powder to fall onto the tape, which is adhered to a substrate, and then invert and gently tap the backside of the substrate to remove the excess powder. The National Physical Laboratory who recently studied the mounting of fibres have produced a protocol for mounting fibres onto silicon wafers using silver dag for SIMS.77 To analyse volatile materials, it is essential to lower their vapour pressures sufficiently by cooling. Polymer surfaces are mobile near or above their glass transition temperature. The surface groups in an aqueous environment can arrange very differently from that in a UHV environment used in XPS and SIMS analysis.5,78 Consequently the surface state in an aqueous environment can be

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maintained by cooling below the glass transition temperature.71 A spectrometer capable of carrying out cryo-analysis would be equipped with cooling in both the introductory and analysis chamber. Condensation of water on the sample surface during cooling can be minimised by cooling the sample in an introductory chamber purged dry-nitrogen or inert gas. Water condensed on the sample surface can be removed by sublimation, meanwhile, SIMS analysis can be performed to find the point at which the surface is free of ice.

8.5

Sources of further information and advice

There are a number of text books describing XPS and ToF SIMS in far more detail than is presented here.79±82 Review articles are constantly appearing for those wishing to find more opinion and comment on applications of surface analysis.83,84 In Table 8.3 a list of relevant conferences, UK meetings and web resources is provided. Table 8.3 Organisations, conferences and internet resources URL

Organisation

http://www.kratosanalytical.net/surface/ XPS manufacturer: Kratos Analytical, UK http://www.specs.com/ XPS manufacturer: Specs, USA http://www.thermo.com/ XPS manufacturer: Thermo VG, UK http://www.phi.com/ XPS and SIMS instrument manufacturer: Physical Electronics, USA. http://www.ion-tof.com/ SIMS instrument manufacturer: IONTOF, Germany http://www.chem.qmul.ac.uk/surfaces/ List of UK, European and worldwide surface sites including academic and commercial equipment and service providers http://www.uwo.ca/ssw/mailinglist.html Popular surface science mailing list for general discussion of all aspects of surface science http://www.ssbii.org.uk/ Surface Science of Biologically Important Interfaces SSBII UK network with annual meetings http://www.uksaf.org/ UK Surface Analysis Forum, biannual UK meetings http://www.avs.org/ AVS meeting, annual US meeting with well attended bio- sessions and plenary events http://www.ecasia.org/ European Conference on Applications of Surface and Interface Analysis, biannual conference

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Acknowledgements

Andrew Hook and Darren Albutt are thanked for their contributions in discussion of this manuscript.

8.7

References

[1] Hearn MJ, Ratner BD, Briggs D. Macromolecules 1988; 21(10): 2950±2959. [2] Hearn MJ, Briggs D, Yoon SC, Ratner BD. Surface and Interface Analysis 1987; 10(8): 384±391. [3] Castner DG, Ratner BD. Surface Science 2002; 500(1±3): 28±60. [4] Ratner BD, Bryant SJ. Annual Review of Biomedical Engineering 2004; 6: 41±75. [5] Ratner BD, Weathersby PK, Hoffman AS, Kelly MA, Scharpen LH. Journal of Applied Polymer Science 1978; 22(3): 643±664. [6] Rawsterne RE, Gough JE, Rutten FJM, Pham NT, Poon WCK, Flitsch SL, Maltman B, Alexander MR, Ulijn RV. Surface and Interface Analysis 2006; 38(11): 1505±1511. [7] Ablett R. Philosophical Magazine 1923; 46(272): 244±256. [8] Bain CD, Whitesides GM. Journal of the American Chemical Society 1988; 110(11): 3665±3666. [9] Whitesides GM, Laibinis PE. Langmuir 1990; 6(1): 87±96. [10] Pertays KM, Thompson GE, Alexander MR. Surface and Interface Analysis 2004; 36(10): 1361±1366. [11] Brewer NJ, Foster TT, Leggett GJ, Alexander MR, McAlpine E. Journal of Physical Chemistry B 2004; 108(15): 4723±4728. [12] Good RJ. Journal of Adhesion Science and Technology 1992; 6(12): 1269±1302. [13] Vohrer U, Blomfield C, Page S, Roberts A. Applied Surface Science 2005; 252(1): 61±65. [14] Tougaard S, Jansson C. Surface and Interface Analysis 1993; 20(13): 1013±1046. [15] Scofield JH. Journal of Electron Spectroscopy and Related Phenomena 1976; 8(2): 129±137. [16] Seah MP, Gilmore IS, Spencer SJ. Journal of Electron Spectroscopy and Related Phenomena 2001; 120(1±3): 93±111. [17] Seah MP, Spencer SJ. Journal of Electron Spectroscopy and Related Phenomena 2006; 151(3): 178±181. [18] Vickerman JC, Gilmore IS, Surface Analysis: The Principle Techniques, 2nd edn, 2009, Chichester, John Wiley & Sons. [19] Briggs D, Beamson G. Analytical Chemistry 1992; 64(15): 1729±1736. [20] Rafati A, Davies MC, Shard AG, Hutton S, Mishra G, Alexander MR. Journal of Controlled Release 2009; 138(1): 40±44. [21] Alexander MR, Payan S, Duc TM. Surface and Interface Analysis 1998; 26(13): 961±973. [22] Alexander MR, Wright PV, Ratner BD. Surface and Interface Analysis 1996; 24(3): 217±220. [23] Hollander A. Surface and Interface Analysis 2004; 36(8): 1023±1026. [24] Clark DT, Abushbak MM, Brennan WJ. Journal of Electron Spectroscopy and Related Phenomena 1982; 28(1): 11±21. [25] http://www.nist.gov/srd/WebGuide/nist71/71imfp.htm. [26] Hall SM, Andrade JD, Ma SM, King RN. Journal of Electron Spectroscopy and Related Phenomena 1979; 17(3): 181±189.

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[27] Cadman P, Evans S, Gossedge G, Thomas JM. Journal of Polymer Science Part C ± Polymer Letters 1978; 16(9): 461±464. [28] Clark DT, Thomas HR, Shuttleworth D. Journal of Polymer Science Part C ± Polymer Letters 1978; 16(9): 465±471. [29] Clark DT, Fok YCT, Roberts GG. Journal of Electron Spectroscopy and Related Phenomena 1981; 22(2): 173±185. [30] Cumpson PJ, Seah MP. Surface and Interface Analysis 1997; 25(6): 430±446. [31] Powell CJ, Jablonski A. Journal of Vacuum Science & Technology A ± Vacuum Surfaces and Films 2001; 19(5): 2604±2611. [32] Yang J, Rose F, Gadegaard N, Alexander MR. Advanced Materials 2009; 21(3): 300±304. [33] Tanuma S, Powell CJ, Penn DR. Surface and Interface Analysis 1991; 17(13): 911±926. [34] Tanuma S, Powell CJ, Penn DR. Surface and Interface Analysis 1991; 17(13): 927±939. [35] Seah MP, Spencer SJ. Surface and Interface Analysis 2002; 33(8): 640±652. [36] Braun RM, Cheng J, Parsonage EE, Moeller J, Winograd N. Analytical Chemistry 2006; 78(24): 8347±8353. [37] Beamson G, Mokarian-Tabari P, Geoghegan M. Journal of Electron Spectroscopy and Related Phenomena 2009; 171(1±3): 57±63. [38] Hinder SJ, Lowe C, Watts JF. Progress in Organic Coatings 2007; 60(3): 255±261. [39] Hinder SJ, Lowe C, Watts JF. Surface and Interface Analysis 2007; 39(6): 467±475. [40] Mahoney CM, Yu JX, Fahey A, Gardella JA. Applied Surface Science 2006; 252(19): 6609±6614. [41] Chen YY, Yu BY, Wang WB, Hsu MF, Lin WC, Lin YC, Jou JH, Shyue JJ. Analytical Chemistry 2008; 80(2): 501±505. [42] Sanada N, Yamamoto A, Oiwa R, Ohashi Y. Surface and Interface Analysis 2004; 36(3): 280±282. [43] Cumpson PJ. Journal of Electron Spectroscopy and Related Phenomena 1995; 73(1): 25±52. [44] http://www.npl.co.uk/nanoscience/surface-nanoanalysis/products-and-services/ arctick. [45] Tougaard S. Surface and Interface Analysis 1998; 26(4): 249±269. [46] Tougaard S, Jablonski A. Surface and Interface Analysis 1997; 25(6): 404±408. [47] Coxon P, Krizek J, Humpherson M, Wardell IRM. Journal of Electron Spectroscopy and Related Phenomena 1990; 52: 821±836. [48] Drummond IW, Street FJ, Ogden LP, Surman DJ. Scanning 1991; 13(2): 149±163. [49] King PL, Browning R, Pianetta P, Lindau I, Keenlyside M, Knapp G. Journal of Vacuum Science & Technology A ± Vacuum Surfaces and Films 1989; 7(6): 3301± 3304. [50] Bullett NA, Short RD, O'Leary T, Beck AJ, Douglas CWI, Cambray-Deakin M, Fletcher IW, Roberts A, Blomfield C. Surface and Interface Analysis 2001; 31(11): 1074±+. [51] Walton J, Fairley N. Surface and Interface Analysis 2004; 36(1): 89±91. [52] Andrade JD, Surface and Interfacial Aspects of Biomedical Polymers, 1985, New York, Plenum Press. [53] Clark DT, Dilks A. Journal of Polymer Science Part A ± Polymer Chemistry 1977; 15(1): 15±30. [54] Clark DT, Dilks A. Journal of Polymer Science Part A ± Polymer Chemistry 1976; 14(3): 533±542. [55] Clark DT, Adams DB, Dilks A, Peeling J, Thomas HR. Journal of Electron Spectroscopy and Related Phenomena 1976; 8(1): 51±60.

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[56] Clark DT, Adams DB. Theoretica Chimica Acta 1975; 39(4): 321±328. [57] Clark DT, Adams DB. Journal of Electron Spectroscopy and Related Phenomena 1975; 7(5): 401±409. [58] Briggs D, Wootton AB. Surface and Interface Analysis 1982; 4(3): 109±115. [59] Benninghoven A, Sichtermann WK. Analytical Chemistry 1978; 50(8): 1180±1184. [60] Wilson RG, Stevie FA, Magee CW, Secondary Ion Mass Spectrometry: A Practical Handbook for Depth Profiling and Bulk Impurity Analysis, 1989, New York, John Wiley & Sons. [61] Briggs D, Hearn MJ. Vacuum 1986; 36(11±12): 1005±1010. [62] Ogaki R, Green F, Li S, Vert M, Alexander MR, Gilmore IS, Davies MC. Applied Surface Science 2006; 252(19): 6797±6800. [63] Briggs D, Brown A, Vickerman JC, Handbook of Static Secondary Ion Mass Spectrometry, 1989, Chichester, John Wiley & Sons Ltd. [64] Davies MC, Short RD, Khan MA, Watts JF, Brown A, Eccles AJ, Humphrey P, Vickerman JC, Vert M. Surface and Interface Analysis 1989; 14(3): 115±120. [65] Chilkoti A, Ratner BD, Briggs D. Surface and Interface Analysis 1992; 18(8): 604±618. [66] Vickerman JC, Gilmore IS, Surface Analysis: The Principle Techniques, 2nd edn. [67] Geladi P, Kowalski BR. Analytica Chimica Acta 1986; 185: 1±17. [68] Wagner MS, Castner DG. Langmuir 2001; 17(15): 4649±4660. [69] Chilkoti A, Schmierer AE, Perezluna V, Ratner BD. Analytical Chemistry 1995; 67(17): 2883±2891. [70] Urquhart AJ, Taylor M, Anderson DG, Langer R, Davies MC, Alexander MR. Analytical Chemistry 2008; 80(1): 135±142. [71] Belu AM, Graham DJ, Castner DG. Biomaterials 2003; 24(21): 3635±3653. [72] Tyler BJ. Applied Surface Science 2006; 252(19): 6875±6882. [73] ISO 18117:2009, Surface chemical analysis ± Handling of specimens prior to analysis. [74] Vickerman JC, Briggs D, ToF SIMS: Surface Analysis by Mass Spectrometry, Chichester, IM Publications, p. 116. [75] Foster TT, Alexander MR, Leggett GJ, McAlpine E. Langmuir 2006; 22(22): 9254± 9259. [76] Yang L, Shirahata N, Saini G, Zhang F, Pei L, Asplund MC, Kurth DG, Ariga K, Sautter K, Nakanishi T, Smentkowski V, Linford MR. Langmuir 2009; 25(10): 5674±5683. [77] Lee JLS, Gilmore IS, Fletcher IW, Seah MP. Applied Surface Science 2008; 255(4): 1560±1563. [78] Senshu K, Yamashita S, Mori H, Ito M, Hirao A, Nakahama S. Langmuir 1999; 15(5): 1754±1762. [79] Briggs D, Seah MP, Practical Surface Analysis: Auger and X-ray Photoelectron Spectroscopy 2nd edn, 1994, Chichester, John Wiley & Sons. [80] Briggs D, J.T. G, Surface Analysis by Auger and X-Ray Photoelectron Spectroscopy, 2003, Chichester, IM Publications. [81] Wolstenholme JF, Watts WJ, An Introduction to Surface Analysis by XPS and AES, 2003, Chichester, John Wiley and Sons. [82] Vickerman JC, Briggs D, ToF SIMS: Surface Analysis by Mass Spectrometry, 2001, Chichester, IM Publications. [83] McArthur SL. Surface and Interface Analysis 2006; 38(11): 1380±1385. [84] Ratner BD. Surface and Interface Analysis 1995; 23(7±8): 521±528.

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Plate III (a) Predicted water contact angle from a PLS model plotted against measured water contact angle. (b) Regression coefficient given by the PLS model showing the low water contact angle associated with polar, hydrophilic ions (labelled ions with negative regression coefficients) and vice versa. Reprinted with permission from Urquhart AJ, Taylor M, Anderson DG, Langer R, Davies MC and Alexander MR, Analytical Chemistry 2008 80(1): 135±142. Copyright American Chemical Society (2008).

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9

Techniques for analyzing biomaterial surface structure, morphology and topography N . S . M U R T H Y , Rutgers ± The State University of New Jersey, USA

Abstract: This is a brief review of the key techniques that are most useful in characterizing the surface structure, morphology, and topography of biomaterials. The emphasis is on the use of these techniques for the evaluation of materials used in prostheses, biomedical implants and tissue scaffolds. The first part describes the techniques using light, electrons and scanning probes to examine the surface morphology. Profilometry that provides quantitative measure of the surface roughness is also presented. Next, techniques that characterize the surface structures at the molecular level, x-ray scattering and Raman spectroscopy, are presented. The enhancements that make these classical techniques surface sensitive are discussed. Finally, contact-angle measurements are discussed in the context of the effect of surface structure and topography on the energetics of the interaction between the substrate and the adsorbing molecules. Principles of each of the techniques are described and illustrated with examples relevant to biomedical applications. Key words: surface structure, microscopy, x-ray scattering, Raman spectroscopy, contact-angle.

9.1

Introduction

Besides surface chemistry, other surface properties such as stiffness and roughness are known to dictate the protein adsorption and the cell response and thus determine the viability of a biomedical device [1]. Surface chemistry provides signals for the cells to attach, proliferate and differentiate. Substrate stiffness affects the cell motility and thus affects the cell behavior. The significance of nanometer to micrometer size topographical features is not well understood, and is still an active area of research. Surface roughness, in particular, influences the wetting behavior and the biocompatibility properties of solid substrates. This in turn affects the performance of biomaterials in various biomedical applications including, sutures, bone pins, implants, and in general, tissue scaffolds. The influence of substrate geometry on cell response has been recognized since the early 1900s. But it is only since the early 1990s that there has been a systematic effort to study the effect of texture on cellular attachment,

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proliferation and motility [2, 3]. Protein adsorption, cellular response and eventually bacterial growth have been shown to be directly affected by surface roughness. These studies have shown that substrate topography in the micrometer range can be used to modify and control the response of the cells and biocompatibility when implanted into tissue. This chapter deals with the characterization of surface topography and structure of materials used in prostheses, biomedical implants and tissue scaffolds. The most common techniques for the physical characterization of surfaces are listed in Table 9.1 [4]. In contrast to the extensive reviews of these techniques that have been published [4, 5], this chapter will present the utility of the most commonly used techniques for examining the surfaces of biomaterials as it relates to protein adsorption and cellular response in biological environments. The goal is not to be exhaustive, but to present recent developments in optical, electron and scanning probe microscopies, x-ray scattering, Raman imaging and surface energy measurements. Principles and the general overview of the different techniques will be presented, and the methods illustrated with examples.

9.2

Surface morphology and topography

Surfaces of biomaterials invariably have features that span length scales from sub-nm to m, and are highly dependent on processing and thermal history. Subnm features are typically due to phase separation and crystallization, and the m-size features are due to steep gradients and pores of different sizes. These surface structures can be discussed in terms of morphology (e.g., domain structure) and topography (e.g., surface roughness due to holes, peaks, ridges and valleys). These features have a large influence on the bulk properties of the material such as strength, but more importantly on the interactions with the surrounding tissue and fluid in a biological environment. It is also possible that surface mechanical properties such as hardness and stiffness influence the surface-protein/cell interactions by themselves, independently of the morphology and surface chemistry. A classical technique for surface hardness measurement is the one based on diamond stylus. Young's modulus can also be calculated from load-displacement measurements. A 2D surface stiffness image can be created by scanning the surface with the stylus. Nanoindentation measurements are routinely performed using scanning force microscopes [6, 7]. These mechanical properties will not be discussed here. Instead, this section will deal with the direct imaging of the surface features on length scales of 1 to 10 m. There are many microscopic techniques that provide surface morphology and topographic data at different length scales. The technique to be chosen depends on the information that is being sought. Some of the factors that need to be considered are the lateral resolution (nm to m), the dimensions of the surface to be examined (m to mm), the depth of the surface (nm to m), and the sample

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Table 9.1 Essential features of the various techniques used for characterizing surface structure and morphology [24]

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Technique

Probe

Depth analyzed (m)

Lateral resolution (m) (magnification)

Information

Comments

Optical microscopy, OM

Light

0.1

0.3 (2 to 2000)

Surface roughness, structure

Many possibilities, good height resolution with interference techniques

Scanning electron microscopy, SEM

Electrons

2

0.1 (20 to 1  105 )

Surface topography

Vacuum technique

Scanning force microscopy, SFM

Cantilever

0.001

0.0005 (1000 to 2  106 )

Surface topography, composition, toughness

Atomic resolution, many different modes

Profilometry

Light/ cantilever

0.01

5

Surface texture

Quantitative measure of surface roughness

X-ray reflectometry, XR Grazing incidence x-ray small-angle scattering, GISAXS

X-rays

0.5

1

Surface roughness, thin surface layers, lateral structure

Flat surfaces required

Neutron reflectometry, NR

Neutrons

0.5

±

Surface roughness, enrichment layer

Deuterated compound needed

Micro-indentation, MI

Cantilever

100

200

Surface hardness, modulus

Quantitative interpretation difficult

Confocal Raman imaging

Light

50

2

Chemical species

Raman active bands required

Pendent drop

Liquid

0.2

Surface tension/ contact angle, ST

Liquid drop

0.1

1000

Interface tension

Indirect technique

Surface energy

Easy to use, molecular information difficult

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environment (vacuum, ambient, metal coated or immersed in water). Optical microscopy is the most versatile of the methods for examining the surface morphology [8]. However, the lateral resolution is limited to ~ 300 nm. Scanning near field techniques (SNOM) and interference techniques enhance the utility of the technique [9]. For higher resolution, the most established techniques are the scanning and transmission electron microscopy, scanning tunneling microscopy and atomic force microscopy.

9.2.1

Confocal microscopy

One of the main drawbacks of a conventional light microscope is that the object planes outside of the focal plane contribute equally to the image as those points at the focal plane, thus producing a blurred image. Confocal laser scanning microscopy (CLSM) [10, 11], also called laser scanning confocal microscopy (LSM), provides a blur-free image by eliminating the out-of-focus light or glare and can be used to collect serial optical sections from thick specimens. In a conventional wide-field microscope, the entire specimen is illuminated from a suitable lamp and the image formed by a series of lenses is viewed directly by eye or captured on a detector. In contrast, in confocal microscopy, the surface is scanned by one or more focused beams of light, usually a laser, and the image is reconstructed on a computer (Fig. 9.1). Laser light from the illuminating aperture passes through an excitation filter (not shown), reflected

9.1 Schematic of the confocal principle in epifluorescence scanning mode.

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by the dichroic mirror and is brought into focus by the objective lens to a diffraction limited spot at the focal plane within the specimen. The light from the specimen is passed through a pinhole that provides the crucial spatial filtering, and is detected by a photomultiplier tube (PMT). Although unstained specimens can be viewed using light reflected from the specimen, the samples are usually stained or labeled with one or more fluorescent probes. In this fluorescence mode, fluorescence emissions excited both within the illuminated in-focus voxel (volume picture element) and within the illuminated cones above and below it, are collected by the objective and pass through the dichroic mirror and the emission filter (not shown). In both of these modes, only the light from the in-focus voxel is able to pass unimpeded through the imaging aperture to be detected by the PMT. The emissions from regions below the focal plane and from above it have different primary image plane foci and are thus severely attenuated by the imaging aperture, contributing little to the final confocal image. The output from the PMT is built into a 2D image. Image resolution is typically 0.2 m in the transverse plane (xy-plane) and ~ 0.5 m in the z-plane. The greatest advantage is the possibility of making a 3D image of the surface of the sample within a depth of 100±200 m. Application Confocal microscopy is widely used to examine surface and near-surface features in biomaterials such as bone [12], dentin and enamel [13], imaging cells on substrates, topography of substrates, and porous structures in scaffolds. Confocal images are typically obtained using the light reflected from the sample (epitransmission or reflectance) or by capturing the fluorescent light that is excited in the sample by the incident beam (epifluorescence). In one example, CLSM was used to study the bimodal porous structure in cancellous bone. The images shown in Fig. 9.2 were obtained in reflectance mode from a single location on defatted and deproteinized cancellous canine bone specimen [12]. Both the surface and near-surface features can be seen in the Z-stack overlay image shown in the figure. The many distinct macropores have micropores along their walls (indicated with arrows). The macropores are interconnected through the micropores in the intertrabecular space. Such features are difficult to see with other imaging techniques. Confocal microscopy is an effective tool to nondestructively image the surface and near-surface structures in a way that is not possible with other techniques such as microcomputed tomography and magnetic resonance imaging.

9.2.2

Scanning electron microscopy (SEM)

SEM reveals the surface features at nm lateral resolutions, about two orders of magnitude better than optical microscopes, and is often preferred over optical

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9.2 Z-stack overlay of the confocal image of a cancellous bone obtained in reflection mode from a Zeiss 210 CLSM (Carl-Zeiss, Jena, Germany) using an Ar 488 nm laser. The figure shows macropores and interconnecting micropores (indicated with arrows) [12]. Reprinted with permission from `Confocal Laser Scanning Microscopy as a Tool for Imaging Cancellous Bone' by I.O. Smith et al., published in Biomed. Mater. Res. Part B: Appl. Biomater. (2006) 79B: 185± 192, ß 2006, Wiley Periodicals, Inc.

microscopy because of the increased depth of field at low resolutions. However, at resolutions approaching sub-m (3±6 nm, best case), SEM is limited by its depth of field (<<1 m at high magnification), and it cannot image subsurface features [12]. SEM uses a beam of electrons that interact with the specimen in a vacuum environment (Fig. 9.3). Various interactions occur between the electron beam and the sample surface including transmission of the primary electrons, secon-

9.3 Schematic diagram of a scanning electron microscope.

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dary electron emission, back scattering, absorption and emission of light and xrays. In transmission electron microscopy (TEM), the transmitted electrons are used to produce an image, the contrast arising from the difference in electron absorption and scattering in different parts of the sample. In SEM, the primary beam is rastered over the sample, and the secondary electrons that are emitted from the surface of the material are collected and composed to provide an electron micrograph. The contrast arises from either electron absorption or emission. Application SEM is widely used because of the fairly simple sample preparation and the ease of using the equipment. It is used to study scaffolds, fiber mats, surfaces of bones and joints. One common application is the use of SEM to examine the surface roughness of implants that can often be related to its performance in vivo. There are several studies in which titanium surfaces have been examined after different types of surface treatments such as after sand-blasting with Al2O3 particles, plasma spraying and elecrolytically coating with hydroxyapatite [14]. Another example of surface treatment is laser ablation. Laser ablation provides a means to modify the surface and thus control the tissue interactions. Figure 9.4 shows an example of how SEM was useful in showing the changes in the surface texture upon ablation of poly(ethylene terephthalate) (PET) surface [15]. PET is widely used in biomedical applications including implantable sutures, surgical mesh, vascular grafts and heart valves [16]. The montage in the figure shows the progressive change in the morphology with the increase in the laser fluence, from an unirradiated area to the left to the ablated area on the right. The figure shows the texture changes from ripple into densely packed cones, before these cones are ablated and removed at higher beam fluence. The figure is most illustrative in that m resolution can be achieved while examining a relatively large area ~ 0.1 mm, and also the large depth of field that can be obtained. Sample size is limited to about 10 mm in diameter. If the sample is not electrically conductive, then a thin (20±30 m) metallic coating, typically gold, platinum or god/palladium alloy is sputter coated onto the sample to prevent electrical charging of the specimen. In addition to imaging the surface features, SEM is often combined with energy dispersive x-ray analysis (EDX) to obtain a map of the elemental distribution.

9.2.3

Scanning force or atomic force microscopy (SFM or AFM)

This is one in a family of scanning probe techniques that have been used successfully in high resolution imaging and molecular imaging of biological

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9.4 A montage of the images from an SEM showing the laser-ablated regions on the surface of PET injection molded plaque. The fluence of the laser beam increases from left to right [15]. Reprinted with permission from: `Self-assembled and etched cones on laser-ablated polymer surfaces' by N.S. Murthy et al., J. Appl. Phys. (2006) 100: 023538, 1±12, ß 2006, American Institute of Physics.

ß Woodhead Publishing Limited, 2011 9.5 (a) Illustration of the principle of AFM. (b±d): A schematic of the three common AFM operation modes. (b) Contact mode. (c) Tapping mode. (d) Phase imaging mode [17]. Reprinted with permission from `Atomic force microscopy of biomaterials surfaces and interfaces' by K.D. Jandt, Surface Science (2001) 491: 303±332, ß 2001, Elsevier.

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macromolecules since their invention in the 1980s [17, 18]. In SFM, the force between a probe and the sample surface is used to map the surface morphology and to study the interaction forces in biological systems down to the picoNewton range, typical of single ligand-receptor interaction. Other techniques, not discussed here, use tunneling current, chemical affinity and other phenomena to probe the surface characteristics. In SFM, a 3D image of the surface is created by scanning a micron-size cantilever across the surface (Fig. 9.5). The cantilever is typically made of silicon or silicon nitride, with a tip whose radius of curvature is ~ 10 nm. The forces between the tip and the surface, which are ~ nN and include contributions from Van der Waals interactions, chemical bonding, electrostatic forces and solvation forces, deflect the cantilever. The deflection of the cantilever due to these forces is measured using a laser spot reflected from the top surface of the cantilever into an array of photodiodes. This signal is used maintain a constant force via a feedback loop, or to maintain a constant height, depending on the scan mode that is chosen. The sample is moved across the tip by piezo crystals. The AFM can be operated in a number of modes depending on the application, and three of these are presented in Figs 9.5(b±d). In the static mode, also called the contact mode, the force between the tip and the surface is kept constant during scanning, and the change in the cantilever deflection required to maintain this constant force is used to obtain an image of the sample's surface topography. This mode is suitable for hard surfaces but not for biomaterials which are soft and whose surface may have weakly adsorbed molecules. In the dynamic mode, also called the tapping mode, the cantilever is vibrated close to its resonant frequency, and the changes in amplitude and phase are monitored. The reduction in amplitude is monitored and used to map the surface topographical features. When the AFM is operated in the tapping mode, it is possible to take advantage of the difference in the phase of the cantilever oscillation and the signal sent to the cantilever. This phase difference depends on the sample's hardness, elasticity and adhesion. This phase lag is monitored and a 2D phase image can be produced to qualitatively map the surface viscoelastic properties of the surface such as to map the distribution of soft biomolecules on a stiff substrate. Application Scanning probe microscopes are useful in characterizing surfaces at atomic resolutions in a variety of environments from ultrahigh vacuum to aqueous solutions. It is also used to study time-dependent phenomena such as conformational change of the molecules and the shapes of the whole cells adsorbed onto a surface, hydration induced changes, crystallization and corrosion processes. A typical contact mode AFM image is shown in Fig. 9.6a, and is compared with a corresponding SEM image of a similar sample in Fig. 9.6b. These images

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9.6 Comparison of (a) AFM (deflection in contact mode) and (b) SEM (secondary electron) imaging technique on similar, dry titanium/titanium oxide surfaces (2  2 m 2 ). Oxide domes are visible in both the images [19]. Reprinted with permission from `Direct observation of hydration of TiO2 on Ti using electrochemical AFM: freely corroding versus potentiostatically held' by J.P. Bearinger, C.A. Orme and J.L. Gilbert, Surface Science (2001) 491: 370± 387, ß 2001, Elsevier.

were a part of a study to characterize the morphology of the titanium surfaces in saline-based oxide hydration events in situ under varying biomedically relevant electrochemical conditions in the hydrated state [19]. The figure shows the surface in air prior to immersion. The oxide domes are more clearly visible in AFM than in SEM. Average dome height in the AFM image is 4:3  1:7 nm and the average dome width is 91  16 nm. Phase imaging is widely used to image soft particles embedded in stiffer matrices. A biologically relevant example of the use of the phase image is the observation of the adsorption of biomolecules in the presence of roughness that is inherent in biomaterial surfaces. AFM images of a National Heart Lung and Blood Institute (NHLBI) reference polydimethylsiloxane (PDMS) with and without adsorbed fibrinogen are shown in Fig. 9.7 [20]. Fibrinogen is a plasma protein important for blood coagulation and platelet aggregation, and PDMS is a commonly used substrate for protein adsorption and cell studies. The images were obtained in the tapping mode. The topographic data is presented in Fig. 9.7a and the phase data in Fig. 9.7b. Proteins were adsorbed at a concentration of 500 ng/ml and the surface concentration is ~ 10 molecules/m2. Although isolated fibrinogen molecules adsorbed onto the surface can be observed in the height image, they are more clearly visible in the phase image in which the topography of the PDSM substrate does not appear because of phase imaging. Highest resolution images are obtained when the AFM is operated in the contact mode. However, because the tip is in constant contact with the sample, the shear forces applied to the sample during scanning can potentially damage

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9.7 AFM tapping mode data from PDMS with and without adsorbed fibrinogen [20]. (a) Height image. (b) Phase image. Reprinted with permission from `Individual plasma proteins detected on rough biomaterials by phase imaging AFM' by N.B. Holland and R.E. Marchant, J. Biomed. Mater. Res. (2000) 51: 307±315, ß 2000, Wiley Periodicals, Inc.

the surface, especially the weakly bound protein molecules adsorbed onto the biomaterials. In the tapping mode, the tip is not in contact with the sample during scanning. Thus the forces applied to the sample surface are negligible. Furthermore, in the tapping mode, phase imaging can be used to map surface inhomogeneities that give rise to variations in viscoelasticity of the material surface but do not alter the physical appearance or the topography of the surface.

9.2.4

Profilometry

The microscopic techniques discussed in Sections 9.2.1±9.2.3 provide an image of the surface texture, but not the quantitative measure of the surface roughness. This can be done by profilometry, in which a probe, mechanical (contact) or optical (noncontact), is passed across the surface [21]. The probe follows the contours at each point on the surface, and the height of the probe at each point is recorded and the resulting 1D scan or a 2D map is analyzed. A mechanical stylus captures the features over a large area (~ 100 mm), with a x-y resolution of 5 m and a z-resolution of 0.01 m. An optical probe such as a confocal laser scanning microscope has a range of ~ 2 mm, x-y resolution of 5 m and the z resolution of 0.01 m. AFM covers a much smaller area (100 m) but provides much higher resolutions (0.2 m in x-y and 1.5 pm in z). The force applied with a mechanical stylus is 5 mN compared to 1 N in AFM. Parameters such as arithmetic average of the absolute values of all points of the profile (Ra), root means square values of all the heights around the mean (Rq) are often used to quantify the roughness. These, however, do not reflect the distance between the features and their shape.

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Application Profilometry is frequently used in the evaluation of surfaces of metallic surgical implants. These implants are treated in various ways to bring about chemical and topographic modifications of the surface to improve the osseointegration process and the biomechanical performance of implants. In one study profilometry was used to measure the surface roughness in three different materials, poly-L-lactic acid (PLLA), poly-DL-lactic acid (PDLA), and sodium alginate hydrogel (AGA-100) used in clinical practice, and correlate this to the osteoblast adhesion. These were coated onto mechanically and chemically treated (sand-blasted and acid-etched) Ti substrates. Figure 9.8 shows how the changes in the surface roughness due to different coatings can be monitored using contact profilometry images. These results were used to identify a specific roughness parameter (peak density) which mainly controls the amount of osteoblast adhesion a crucial factor in rapid `osteointegration' [22]. The average roughness was 4.31 in the control (sand-blasted and acid-etched) sample, and was 2.39, 1.39 and 2.78 in AGA-100, PLLA and PDLA coated samples,

9.8 Contact profilometry of Ti surfaces [22]. (a) The initial sand-blasted and acid-etched surface prior to coating. (b) Coated with PLLA. (c) Coated with PDLA. (d) Coated with sodium alginate hydrogel AGA-100. Reprinted with permission from `Contact profilometry and correspondence analysis to correlate surface properties and cell adhesion in vitro of uncoated and coated Ti and Ti6Al4V disks' by A. Bagno et al., Biomaterials (2004) 25: 2437±2445, ß 2001, Elsevier.

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respectively, reflecting the modification of the surface roughness to different degrees depending on the coating material. Only AGA-100 was able to preserve, to some extent, the roughness in the starting surface. Only the starting and AGA100 coated samples allowed the osteoblasts to adhere, and the other two seemed to hinder the adherence. Further evidence suggests that it is indeed this surface roughness and not the chemical composition that is responsible for the differences in adhesion. Profilometers generate an image of the surface height. Size of the area measured and the size of the probe set the upper and lower limits on the size of the features that can be characterized. The nature of the probe and the force exerted by the probe on the surface limit the range of surfaces that can be investigated by these techniques. Although optical techniques are more appropriate for relatively soft materials, stylus technique is more appropriate for examining large areas with substantial slopes within the surface structure [23]. The advantages of this method are that it is direct, less expensive and reproducible.

9.3

Surface structure and spatial distribution

The techniques described in the previous section are useful in examining the morphology, texture and surface topography. At the other extreme, there are surface spectroscopic techniques such as Auger electron spectroscopy, x-ray photoelectron spectroscopy and secondary ion mass spectrometry that provide the chemical makeup or the elemental composition at or near the surface. These techniques are discussed in Chapter 8. Between these two extremes, there is another category of techniques for examining the structure, the molecular organization, surface orientation that change as a result of surface treatment. These structures near the surface can be studied using methods that are surfaceenhanced versions of the classic x-ray scattering, IR and Raman spectroscopic techniques. These classical techniques have been adapted so that x-ray, IR or visible light penetrate only the surface layers of the material, and provide information about the surface structure. This section will discuss the use of x-ray scattering and Raman spectroscopic methods for analyzing the surface structures that arise as a result of surface orientation, phase separation, crystallization and degradation.

9.3.1

X-ray reflectivity and scattering

When x-rays are incident on a surface at very shallow angles, two events can occur. First, just like light, x-rays are also reflected from surfaces. But, because the refractive index of solids for x-rays is < 1, this specular reflection (reflected angle equal to incident angle) occurs for incidence angles less than the critical angle, ~ 0.1ë (Fig. 9.9a). In this configuration, known as reflectometry, the profile of the reflected beam is analyzed to characterize roughness as well as investigate phenomena such as interdifussion, blending, surface-induced order,

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adsorption or surface enrichment of components [24]. X-ray reflectometry is an ideal tool to quantify the surface roughness in the nm length scale. If the interface between the sample and air is not perfectly sharp and smooth then the reflected intensity will deviate from that predicted by the simple Fresnel equations. These deviations are analyzed to obtain the density profile of the interface normal to the surface. Such reflectivity measurements provide surface roughness [25] and the structure of deposited membranes [26]. In a related second technique, the x-rays are incident at a very shallow angle (, Fig. 9.9b) such that the path within the sample is considerably long so as to attenuate the beam within a small depth from the surface. This depth depends on the absorption coefficient of the substrate, and typically a few m for polymers. In this technique, called grazing-incidence scattering, the intensity measured function of the scattering angle 2 will reveal the structure near the surface of

9.9 (a) Schematic of the sample geometry used in x-ray reflectivity measurements. (b) Geometry of the sample, incident and scattered beam used in the data collection in grazing-incidence scattering. (c) Measured x-ray reflectivity plotted as R/RFresnel versus Qz, where RFresnel is the reflectivity of an infinitely sharp, step-like interface [26]. Reprinted with permission from `Characterization of biological thin films at the solid-liquid interface' by C.E. Miller et al., Phys. Rev. Lett. (2005) 94: 238104, 1±4, ß 2005, American Institute of Physics. (d) Arrangement of the head groups and the hydrocarbon tail region modeled based on the data in (c).

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the sample. By varying the angle of incidence, one can probe layers of increasing thickness and analyze the structure and chemical composition by means of diffraction or fluorescence experiments. Grazing incidence small-angle x-ray scattering (GISAXS) provides surface sensitive information at length scales of 1±100 nm. Such measurements have been used to tailor the island density, shape and size distributions. Applications The example shown in Fig. 9.9 illustrates the use of x-ray reflectivity to characterize the phospholipids bilayer membranes at the solid-water interface. The x-ray reflectivity curve typically obtained is shown in Fig. 9.9c. This curve is modeled using a model, in this case, a 4-slab model made of outer head groups, hydrocarbon tails, inner head groups, and water cushion (Fig. 9.9d). The lipid used in this case was 1,2-Dioleoyl-Sn-Glycero-3-phosphocholine (DOPC). Ê hydrocarbon tail region, outer head The data can be explained in terms of 23.2 A Ê Ê group region thickness of 10 A with a 6 A roughness, and the inner head group Ê with 3.8 A Ê roughness. Finally, the model also suggests a thin thickness of 8 A Ê 4 A layer of water between the lipid and the quartz substrate. Such x-ray reflectivity measurements have been used to model the surface roughness as well as the structure near the solid-solution interface, and in the case just discussed, that of the lipid bilayer. The technique requires that the surfaces be flat, extending to a few centimeters. The measurements are performed with x-rays illuminating the surface at grazing incidence, typically a few milliradians close to the critical angle for total reflection at the material-air interface. Because of the reduced intensity due to tight collimation, such measurements are typically performed at synchrotron sources.

9.3.2

Confocal Raman spectroscopy

Raman spectroscopy, like infrared spectroscopy, is highly sensitive to molecular species present in the material. Combining this with the confocal principle described in Section 9.1.2, makes it a powerful tool for obtaining the structural information near the surface. Unlike infrared spectroscopy, Raman spectroscopy is relatively insensitive to water, and hence is potentially more useful in the characterization of the microstructure of biomaterials where water is invariably present. Recent advances in laser sources, optical elements and detector technology have enabled mapping of the surfaces based on the identity and concentration of the selected species. It can be used to nondestructively image the surface to a depth of ~ 50 m, and has been used, for instance, to analyze the surface and assess the residual stress fields in artificial hip joints [27, 28], in addition to studying the oxidation state, crystallization behavior and phase fractions.

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In typical analytical measurements, the Raman spectral intensities are directly proportional to the concentration of the corresponding species within the relatively small deviation in Raman response. Concentrations can be measured from the intensities of an interference-free peak after calibration using a reference material. The most common way to Raman map a surface is to find a characteristic strong band in the species of interest, a band that is free from interference from other species, and map the intensity of this band. However, when strongly overlapping bands are present, as in most biomaterials, chemometrics methods such as multivariate curve resolution are used to obtain the Raman images [29]. Spatial resolutions of 2 m in the x±y plane and 3 m to a depth of 50 m in the z-plane are achievable. Application Confocal Raman imaging is commonly used to examine the changes that occur near the surface due to adsorption of molecules, hydration (with D2O labeled samples), and degradation. By combining the micrometric lateral resolution of a laser beam focused on the polymer surface with a proper selection of Raman bands, it is possible to obtain a highly resolved 2D map of the conformational population patterns, including crystalline and amorphous phase fractions, and oxidation states. In one example, confocal spectroscopic techniques were used to study the microscopic features due to wear in acetabular cups made of ultra-high molecular weight polyethylene (UHMWPE) before and after implantation in vivo (Plate IV between pages 208 and 209) [30]. The figure shows significant oxidation gradients along the subsurface of the long-term implanted retrieval, especially in the main wear zone. A clear relationship is found between orthorhombic crystalline fraction and degree of oxidation along the subsurface, and this was attributed to partial crystallization of polyethylene on oxidation. It is important to note that these confocal experiments show relatively high crystallinity in the immediate subsurface of the cup during implantation lifetime in vivo. This could be related to the formation of polyethylene debris on wear contact. Thus, near-surface UHMWPE gets oxidized and crystallizes during aging in vivo, which could induce significant embrittlement of the polymeric subsurface.

9.4

Energetics

Surface energy is the work required to increase unit surface area at constant temperature, pressure and composition. Surface tension, which is related to surface energy, determines the wettability of a surface and as such plays an important role in determining the biocompatibility, adsorption processes and adhesion in biomaterials. It reflects surface roughness, and is sensitive to both

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surface segregation of components and surface contamination. Surface tension also reflects the strength of bonding within the bulk materials. Hard solids with covalent, ionic or metallic bonds have high surface energy surfaces (surface tension 500±5000 mJ/m2). Most polymers are soft solids in which van der Waals and hydrogen bonding forces play a major role and which have low energy surfaces (< 100 mJ/m2).

9.4.1

Contact angle measurements

The surface tension is most often determined by measuring the contact angle of a drop of liquid positioned on the surface. The contact angle is the angle at which a liquid-vapor interface meets the solid surface. This angle is determined by the interactions across the three interfaces. In Fig. 9.10, SL, SV and LV are the interfacial energy (surface tension) at the solid-liquid, solid-vapor and liquid-vapor interfaces, respectively, and the Young's relation given in the figure relates the contact angle to these three forces. There are several methods of determining the surface tensions including measuring the direct force by Wilhelmy balance, contact angles using sessile drop or captive air bubble method, and capillary penetration into porous systems [31]. Contact angle using sessile drop is the most widely used and will be illustrated here. A drop of fluid, water for most biomaterials, is placed on the surface of interest. The contact angle is determined from the tangent to the liquid drop at contact where the liquid and the solid intersect. Surface roughness or heterogeneity can be evaluated by measuring the contact-angle hysteresis, i.e. by measuring the advancing and receding contact angles [32]. As the volume of drop is increased, and the drop advances over the surface, the edge of the liquid is pinned at the boundaries of the wetting and non-wetting interfaces (if the surface is heterogeneous) or the sharp edge of the surface boundary (if the surface is rough). Thus, the contact angle will be higher than the equilibrium value. Similarly, because of the same pinning the contact angle will be lower than the equilibrium value when the volume of the droplet decreases as the liquid is withdrawn from the droplet. The resulting hysteresis can be used to characterize the surface roughness or heterogeneity.

9.10 Illustration of the Young's force balance giving the equilibrium contact angle.

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Application The use of contact-angle measurement will be illustrated with studies of dynamic contact-angle measurements and adsorption of -globulin made as a function of Wenzel ratio (roughness factor calculated as the ratio of the actual and the projected surface areas) [33]. Contact-angle hysteresis measurements were made on well-defined nanostructured surfaces that were fabricated by growing germanium nanopyramids on Si(001) surfaces. AFM images shown in Fig. 9.11 indicate the different surface morphologies. These surfaces are made of nanopyramids and the density of these pyramids determines the surface roughness It has been reported that surface roughness modifies the contact angles and the contact-angle hysteresis of wetting. There is no consensus as to whether the contact angle increases with roughness and on the effect of roughness on the hysteresis behavior [33]. The results show that the advancing contact angle of water monotonically increases by 20ë from the flat substrates to substrates with maximum pyramid density whereas the receding contact angle remains constant. The biocompatibility of materials is determined at the most fundamental level by protein adsorption. The amount of adsorbed proteins bovine -globulin (BGG) increases significantly with the density of nanopyramids on the substrate, i.e., with the roughness of the surface. Increase in the effective surface by 7% causes a 2- to 4-fold increase in the adsorbed protein. However, the activity of the adsorbed BGG was found to decrease with pyramid density. These results demonstrate that wetting behavior and biocompatibility are both strongly correlated with the surface nanoarchitecture, and thus illustrate the importance of careful surface topographic measurements in understanding the surface-induced biological processes.

9.5

Future trends

With great advances in surface modification discussed elsewhere in this book, the need for rapid characterization of the surface features inexpensively and reproducibly will continue to grow. Optical microscopy will continue to be a most versatile tool to examine surface morphology. Moving beyond the confocal microscopes, near-field scanning optical microscopy can deliver resolutions normally associated with electron microscopy [34]. It uses tapered glass optical elements and techniques of scanned probe imaging (e.g., AFM) to provide images at 50 nm resolution, far below the Rayleigh resolution limit of half the wavelength of the imaging radiation (~ 250 nm). A major limitation of SEM is that the samples that are imaged are typically in a high vacuum environment. Innovative sample chambers are now being used that permit the samples to be in different environments such as moisture. One trend in SEM has been to work at a low incident beam energy (< 5 keV) in order to minimize beam penetration,

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9.11 (a±d) AFM images indicating the increase in surface roughness with increasing nanopyramid density. (e) Advancing (open circles) and receding (filled circles) contact angle hysteresis for water in surfaces with different nanopyramid density. (f) Protein adsorption of BCG and anti BCG vs. roughness factor [34]. Reprinted with permission from `Impact of nanometerscale roughness on contact-angle hysteresis and globulin adsorption' by B. Muller, et al., J. Vac. Sci. Technol. (2001) 19: 1715±1720, ß 2001, AVS: Science & Technology.

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9.11 Continued

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beam spreading, and sample charging and thus improve surface image resolution. Alternatively, high energy beams are being used obtain complementary information such as subsurface features that can be combined with the data obtained at low energies. Scanning probe techniques continue to evolve with the development of new probes to map different aspects of the surface. Availability of high-intensity x-ray synchrotron sources makes it much easier to obtain reflectivity and grazing incidence measurements for the detailed evaluation of surface structures. In more traditional techniques such as contact angle measurement, automation of the measurement and analysis enables highly reliable data to be obtained even by those who may lack specialized skills. In all these techniques, technology is used to move towards the ultimate goal of viewing features at high resolutions in a large dynamic system at high imaging speeds.

9.6

References

1. Wong, J.Y., J.B. Leach and X.Q. Brown, Balance of chemistry, topography, and mechanics at the cell±biomaterial interface: issues and challenges for assessing the role of substrate mechanics on cell response. Surface Science, 2004. 570: 119±133. 2. Chen, C.S. et al., Geometric control of cell life and death. Science, 1997. 276: 1425± 1428. 3. Huang, S. and D.E. Ingber, Shape-dependent control of cell growth, differentiation, and apoptosis: switching between attractors in cell regulatory networks. Exp. Cell. Res., 2000. 261: 91±103. 4. Stamm, M., ed. Polymer surfaces and interfaces: characterization, modification and applications, 2008. Berlin: Springer. 5. Merrett, K. et al., Surface analysis methods for characterizing polymeric biomaterials. J. Biomater. Sci. Polymer Edn, 2002. 13(6): 593±621. 6. Haque, F., Application of nanoindentation to development of biomedical materials. Surface Engineering, 2003. 19(4): 253±268. 7. Bedoui, F., F. Sansoz and N.S. Murthy, Incidence of nanoscale heterogeneity on the nanoindentation of a semicrystalline polymer: experiments and modeling. Acta Materialia, 2008. 56: 2296±2306. 8. Sawyer, L.C. and D.T. Grubb, Polymer Microscopy. 3rd edn, 2008. Springer, 540. 9. Ohtsu, M. and H. Hori, Near-field nano-optics ± From basic principles to nanofabrication and nanophotonics, 1999. New York: Kluwer Academic/Plenum. 10. Semwogerere, D. and E.R. Weeks, 'Confocal microscopy', in Encyclopedia of Biomaterials and Biomedical Engineering, 2005. pp. 1±10. 11. Shotton, D., Confocal scanning optical microscopy and its applications for biological specimens. Journal of Cell Science, 1989. 94: 175±206. 12. Smith, I.O. et al., Confocal laser scanning microscopy as a tool for imaging cancelous bone. J. Biomed. Mat. Res. Part B: Appl. Biomat., 2006. 79B: 185±192. 13. Curtis, R.V. and T.F. Watson, eds. Dental biomaterials: Imaging, testing and modelling, 2009. Cambridge: Woodhead Publishing. 14. Prado da Silva, M.H. et al., Surface analysis of titanium dental implants with different topographies. Mat. Res., 2000. 3(3): 61±67. 15. Murthy, N.S. et al., Self-assembled and etched cones on laser-ablated polymer surfaces. J. Appl. Phys., 2006. 100: 1±12.

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16. von Recum, A.F., ed. Handbook of Biomaterials Evaluation: Scientific, Technical and Clinical Testing of Implant Materials, 2nd edn, 1998. Boca Raton, FL: CRC Press. 17. Jandt, K.D., Atomic force microscopy of biomaterials surfaces and interfaces. Surface Science, 2001. 491(3): 303±332. 18. Siedlecki, C.A. and R.E. Marchant, Atomic force microscopy for characterization of the biomaterials interface. Biomaterials, 1998. 19(4±5): 441±454. 19. Bearinger, J.P., C.A. Orme and J.L. Gilbert, Direct observation of hydration of TiO2 on Ti using electrochemical AFM: freely corroding versus potentiostatically held. Surface Science, 2001. 491: 370±387. 20. Holland, N.B. and R.E. Marchant, Individual plasma proteins detected on rough biomaterials by phase imaging AFM. J. Biomed. Mater. Res., 2000. 51: 307±315. 21. Assender, H., V. Bliznyuk and K. Porfyrakis, How surface topography relates to materials' properties. Science, 2002. 297: 973±976. 22. Bagno, A. et al., Contact profilometry and correspondence analysis to correlate surface properties and cell adhesion in vitro of uncoated and coated Ti and Ti6Al4V disks. Biomaterials, 2004. 25: 2437±2445. 23. Wennerberg, A. et al., Characterizing three-dimensional topography of engineering and biomaterial surfaces by confocal laser scanning and stylus techniques. Med. Eng. Phys., 1996. 18: 548±556. 24. Stamm, M., Polymer surfaces, interfaces and thin films studied by x-ray and neutron reflectometry, in Scattering in polymeric and colloidal systems, W. Brown and K. Mortensen, eds, 2000. Amsterdam: Gordon and Breach Science Publishers, pp. 495±534. 25. Stone, V.W. et al., Roughness of free surfaces of bulk amorphous polymers as studied by x-ray surface scattering and atomic force microscopy. Phys. Rev. B, 1999. 60: 5883±5894. 26. Miller, C.E. et al., Characterization of biological thin films at the solid-liquid interface by X-ray reflectivity. Phys. Rev. Lett., 2005. 94: 238104 (1-4). 27. Pezzotti, G., Stress microscopy and confocal Raman imaging of load-bearing surfaces in artificial hip joints. Exp. Rev. Med. Dev., 2007. 4(2): 165±189. 28. Pezzotti, G. et al., Confocal Raman spectroscopic analysis of ceramic hip joints. Key Engineering Materials, 2006. 309±311: 1211±1214. 29. Andrew, J.J. and T.M. Hancewicz, Rapid analysis of Raman image data using twoway multivariate curve resolution. Appl. Spectro., 1998. 52: 797±807. 30. Pezzotti, G. et al., Confocal Raman spectroscopic analysis of cross-linked ultra-high molecular weight polyethylene for application in artificial hip joints. J. Biomed. Opt., 2007. 12(1): 0141011 (1±14). 31. Grundke, K., Characterization of polymer surfaces by wetting and electrokinetic measurements ± contact angle, interfacial tension and zeta potential, in Polymer surfaces and interfaces, M. Stamm, ed., 2008. Berlin: Springer. 32. Extrand, C.W., in Encyclopedia of surface and colloid science, P. Somasundaran, ed., 2006. Boca Raton, FL: CRC Press, pp. 2876±2891. 33. Muller, B. et al., Impact of nanometer-scale roughness on contact-angle hysteresis and globulin adsorption. J. Vac. Sci. Technol., 2001. 19(5): 1715±1720. 34. Doose, S., Trends in Biological Optical Microscopy. Chem. Phys. Chem., 2008. 9: 523±528.

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Plate IV (a) Photograph of a long-term retrieved cup. (b±e) Maps of wear index (related to oxidation index) collected at increasing depths of focal plane z0 of the confocal probe on the long-term retrieved UHMWPE acetabular cup shown in (a). (b and d): Nonwear zone. (c and e): Wear zone. Reprinted with permission from Pezzotti G et al., `Confocal Raman spectroscopic analysis of cross-linked ultra-high molecular weight polyethylene for application in artificial hip joints', J. Biomed. Opt. (2007) 12: 014011 1±14, 2007, SPIE.

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Modifying biomaterial surfaces to optimise interactions with blood A . D E M E L , Y . R A F I E I and B . G . C O U S I N S , University College London, UK and A . M . S E I F A L I A N , University College London, UK and Royal Free Hampstead NHS Trust Hospital, UK

Abstract: This chapter discusses the surface modification of biomaterials designed to interact directly with the blood and surrounding vasculature. In particular, it focuses upon the management of cardiovascular disease through direct surgical intervention and the application of medical devices that are necessary to perform life-saving treatments. Key words: cardiovascular disease, blood-biomaterial interactions, intimal hyperplasia, vascular bypass graft, endothelialisation.

10.1

Introduction

Cardiovascular and coronary artery disease are responsible for a significant percentage of mortality and morbidity in the ageing population, and is expected to increase in the coming years.1 Vascular occlusive disease carries the greatest risk factor, especially in terms of its impact in the coronary arteries where cardiac ischemia may lead to complete heart failure. The main options for reperfusionbased surgical intervention in the treatment of these diseases include angioplasty, stenting, endarterectomy and bypass graft surgery, depending on the degree of occlusion. In cases where the proportion of occluded arteries is greater than 70%, the condition has to be treated with bypass grafts. However, in cases where occlusion rates are less than 70%, percutaneous transluminal coronary angioplasty (PTCA) is employed. This treatment is also sometimes used as a nonsurgical, lower risk, alternative option to thrombolysis. Although drug-eluting stents presented greater initial benefits, their association with the subsequent development of thrombosis2 has led researchers to search for better alternatives.3±6 Autologous bypass conduits are preferred for primary revascularisation in the case of small diameter bypass grafts.7 However, 3±30% of patients are presented with no autologous vessels, as a result of previous disease conditions, and thus there is a need for vascular grafts that are capable of performing in a similar way to autologous vessels.8,9 Heart valve replacement is a standard

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surgical procedure in the treatment of severe valvular heart disease, which can occur as stenosis. Graft thrombogenicity due to material surface incompatibility, and altered flow dynamics at the site of anastomosis or distal outflow, are recognised as primary reasons for blood-contacting device faliure.10±12 Vascular bypass grafts, stents, heart valves and dialysis access devices are now in high demand. There is an urgent clinical need for improved cardiovascular devices which promote desirable blood-biomaterial, cellular and tissue interactions. Recent strategies focus upon surface techniques by modifying the physicochemical properties at the implant surface, and by combining a biomimetic approach through functionalisation, present an exciting challenge to search for ways to improve patency. In this chapter we focus on the surface modification of cardiovascular biomaterials used in the development of prosthetic bypass graft materials, such as expanded poly(tetrafluoroethylene) (ePTFE), poly(ethylene tetrapthalate) (Dacron) and polyurethane (PU), and discuss their applications in the production of blood-contacting devices.

10.1.1 Events at the blood-biomaterial interface When positioning a medical device to repair damage or arrest disease, the initial incision of the tissue causes localised trauma and injury to the blood vessels. Biomaterials that are exposed to the blood and surrounding tissue fluid adsorb specific proteins onto the surface, and restructure the blood-biomaterial interface. This event happens very rapidly (taking place in less than 1 second), and leads to the formation of a thin protein film of a thickness in the order of nanometres.13,14 The adsorption of proteins (composed of polar, non-polar, and charged side groups) contributes to the surface activity. Once present at the surface, protein molecules interact with water, electrolytes and the underlying surface chemistry (and energy) of the material through hydrogen bonding, van der Waals, pi-pi (-) stacking, and electrostatic interactions. Exactly which force governs the interaction of proteins on surfaces depends upon the particular protein, as well as other factors including size, charge, conformation and unfolding rate, as described by Vroman.15 Chemical and physical properties of the materials, such as surface chemistry, energy and topography, influence the interfacial behaviour adjacent to the biomaterial. The interfacial region at the blood-biomaterial surface continually alters and redistributes the protein/ electrolyte/water layer, and the host cells and tissues react to changes in this layer.16 The adsorption of plasma and extracellular matrix (ECM) proteins (fibrinogen, albumin, and -globulin) and, to a lesser degree, fibronectin, collagen, von Willebrand factor (vWF), coagulation factors XI and XII, and high molecular weight kininogen (HMWK), play a crucial role in balancing thrombosis and haemostasis.17,18 Such proteins direct and aid the adhesion of red blood cells, platelets (the first cellular components to adsorb to the protein film), then

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leukocytes, and endothelial cells (EC). The cellular components interact with the protein layer to guide migration, initiate blood coagulation, and stimulate cell proliferation and differentiation, as specific proteins present binding sites for macromolecules and receptors, guiding the recruitment of further cells interacting within the vasculature.

10.1.2 Blood coagulation The initial events leading to thrombosis surrounding the tissue-implant interface are mediated by surface interactions with adsorbed proteins (intrinsic pathway) or through the release of tissue factor (TF) from damaged cells at the site of injury (extrinsic pathway).19 The intrinsic pathway is independent of injury. Adsorbed surface proteins form a complex composed of collagen, HMWK, prekallikrein, and factor XII.19 Inactive precursors (clotting factors) change conformation and are converted into active enzymes via a biochemical cascade, which, with the addition of certain co-factors, results in platelet activation. The cleavage of prothrombin via the prothrombinase complex bound to cellular membranes, generates thrombin and by converting fibrinogen to fibrin, forms a stable insoluble gel (red thrombus or clot).19 Vascular injury and damage to the endothelium releases TF, collagen and vWF to initiate the extrinsic pathway. Clotting factors interact with platelet surface receptors and play a fundamental role in the interaction of collagen to initiate thrombosis, release growth factors and cytokines to enhance the coagulation cascade and strengthen the haemostatic plug. The platelets change morphology and agglomerate, forming a thrombus layer. It is important to note that both pathways converge during the formation of the prothrombinase complex, leading to the generation of the thrombin referred to as the common pathway. However, the events described here are simplified, and the reader is encouraged to locate further examples from the excellent reviews that have been produced in this area.19±21

10.1.3 Thrombosis and thrombogenicity Vascular procedures, such as arteriovenous graft placement and angioplasty, damage the adventitial and medial tissues of the arterial wall, causing injury to the endothelium lining the intima.22 Angioplasty, for example, is a controlled traumatic event, which is aimed at causing plaque rupture by the widening of a narrowed or obstructed vessel. These processes can expose the otherwise intact sub-endothelial matrix, removing the protective endothelium and exposing the medial smooth muscle cells (SMC) directly to blood flow, other procoagulants and proinflammatory blood constituents. Tissue trauma rapidly initiates the recruitment of inflammatory cells, which release potent cytokines and promote the migration and proliferation of SMCs.23 The anti-coagulant and vascular

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protective functions performed by prostacyclin (PGI2) and nitric oxide (NO) in the intact endothelium, which are required for the regulation of blood flow, soon diminish.23 Both molecules are necessary to inhibit platelet adhesion, aggregation and activation in the endothelium and SMC, which are considered to be early events in the development of intimal hyperplasia (IH).23* Furthermore, NO inhibits SMC proliferation and migration. In addition, the adventitial layer is partially removed for the creation of the anastomosis during surgery, which deprives the vessel wall of oxygen and vital nutrients.23,24 Prosthetic bypass grafts, which are not mechanically or biologically optimised, can further aggravate thrombogenic incidences due to the lack of a fully functioning endothelium. Both IH and thrombosis are the principal mechanisms responsible for cardiovascular graft failure, which can trigger myocardial infarction, or occlude smaller arteries when emboli become detached from the surface. Thrombogenic materials activate the immune and complement systems to stimulate further cell recruitment and macrophage fusion to form foreign body giant cells (FBGC) during the development of a fibrous tissue capsule surrounding the graft material. Fibrous encapsulation (the thickness of which can be considered to be one measure of thrombogenicity) can lead to the failure of the implant, which can often be treated only by further surgery and its removal. Thrombogenicity refers to the ability of a material to produce a thrombus (clot), and is defined as the ability to induce or promote the formation of thromboemboli.25 It is known that low rates of thrombosis can be tolerated since the fibrinolytic system exists to break down fibrin during clot lysis. The criterion for non- or anti-thrombogenic materials was originally thought to represent long clotting times followed by minimal platelet deposition.21 However, this has been extended to define desirable parameters when considering anti-thrombogenicity: (1) A low thrombin production rate constant (kp <10ÿ4 cm s1), (2) low platelet consumption and activation, (3) some platelet adhesion, and (4) low complement and leukocyte activation.21 Such criteria highlight essential parameters that identify a strategy for producing anti-thrombogenic biomaterials.21 However, there is shortage of direct and simple laboratory-based test methods to measure and evaluate the thrombogenic behaviour of biomaterials used for cardiovascular applications. Almost all materials are considered to be thrombogenic, with the exception of EC, which lines the vasculature. Large diameter vascular grafts were originally thought to be anti-thrombogenic in nature. For example, ePTFE bypass grafts appear non-thrombogenic due to the high flow rates of blood past the luminal surface; but in reality, they all are thrombogenic to a certain degree.

*

Intimal hyperplasia ± a new or thickened layer of arterial intima formed on a prosthesis (or in atherosclerosis) by the migration and proliferation of cells within an organ or tissue beyond that which is ordinarily seen.

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10.1.4 Blood flow (shear stress) In healthy individuals the flow of blood is laminar. In comparison, the flow of blood in diseased or occluded arteries may often be transitional or even turbulent in behaviour. At the blood-biomaterial interface, haemodynamic forces of shear stress at the wall surface play a critical role in blood-contacting devices, and influence protein adsorption,26 platelet and leukocyte adhesion.27 Leukocytes recognise specific proteins and adhere under flowing conditions to initiate further cell signalling and recruitment events. A study evaluating leukocyte adhesion on PU materials has shown that cell density decreased with increasing shear stress.28 Certain shear stress models have been studied (specifically, when applied to seeded vascular grafts) to promote EC retention29±31 and have been found to correlate with changes in the EC phenotype.32

10.1.5 Cell-material interactions Cell-material interactions can be divided in to the sub-categories of ligandreceptor and non-receptor mediated interactions (which will be discussed in greater detail later). Ligands are defined herewith as biomolecules that interact with receptors such as integrins. Their geometric spatial arrangement,33±35 density,36,37 orientation, conformation,38 as well as stereochemistry of the sequence,39 all affect cell behaviour. For example, studies have shown that EC have a greater affinity for immobilised cyclic RGD (Arg-Gly-Asp) compared to linear RGD.40 The cyclic peptide is believed to closely resemble the conformation of the native ligand. Rapid cell recruitment and adhesion has been observed with cyclic RGD and would prove beneficial for in-situ endothelialisation studies. A series of cyclic RGD peptides have also been shown to act as potent selective antagonists for the platelet integrin 2 3 (GPIIb/IIIa) and to inhibit platelet-mediated thrombus formation.41 Water-soluble low molecular weight RGD containing peptides interact with integrins and inhibit cell adhesion.42 Furthermore, RGD spatial clustering is also known to be important in determining EC adhesion and the strength of that adhesion.43

10.1.6 Cell-receptor mediated interactions Integrins represent a large family of heterodimeric, transmembrane, cell surface proteins that govern adhesive interactions between cells and macromolecular components of the ECM via binding sequences, such as the RGD motif,44 which binds to V 3 integrin.45,46 Binding to integrin receptors activates signaling transduction pathways that mediate cell attachment and proliferation, the organisation of cytoskeletal actin and the formation of focal adhesions.40,47 However, many integrin receptors are not ligand specific and have an affinity towards multiple isoforms (or combinations).48 Currently EC have been shown

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to express at least 13 different integrins, depending on their state of development, differentiation and function.49 The migration of endothelial progenitor cells (EPC) to ischemic tissues is believed to involve 4 1 integrins, whilst 1, V 3 and V 5 are involved in mediating the adhesion of progenitor cells during endothelialisation.50 Furthermore 2 has shown a significant role in EPC homing to sites of ischemia and in neo-vascularisation.51 Interestingly, 4 integrins are absent in platelets, and ligands aimed at interacting with these receptors may be beneficial in the introduction of in-situ endothelialisation of vascular grafts.52 Apart from integrins, c-kit is another receptor expressed by EPC that may be important for in-situ endothelialisation. Stem cell factor (SCF) is a transmembrane protein which is a ligand for c-kit, and has been shown to be involved in EPC adhesion to the endothelium.53,54 CXCR2 is another receptor expressed in EPC, which plays a key role in the homing of EPC to sites of injury during revascularisation.55

10.1.7 Steps toward endothelialisation of biomaterials Current bioengineering strategies aim to incorporate a fully functional EC layer on the luminal surface of blood-contacting materials.56±58 Studies involving in vitro endothelialisation with cultured EC prior to implantation have shown that the cell layer prevents thrombogenic complications and improves longer-term patency rates.59±62 However, such techniques remain labour intensive and costly, and require a considerable amount of time and expertise, all of which factors mean this process is generally limited to specialised centres. The endothelialisation of synthetic graft surfaces does not occur spontaneously on non-passivated materials, and graft seeding has received a great deal of attention. There are currently two strategies for seeding biomaterials and devices with EC: (1) is a two-step procedure, and (2) a single-step process.63 In twostage seeding, the EC are extracted and cultured for a period of 2±4 weeks in a laboratory in order to expand and culture the cells at high density before seeding. In contrast, the single-stage procedure involves harvesting the cells and culturing them in direct contact with the graft surface.64 EC are usually extracted from veins, arteries, and subcutaneous fat, as these tissues are easily accessible in clinical practice.65 EPC has been extracted from peripheral blood, before being cultured and seeded on decellularised tissue (porcine iliac vessels).66 Recent investigations have focused on the use of stem cells derived from adipose tissue for the tissue engineering of bypass grafts 67 and there have been reports of seeded grafts moving towards clinical trial.68 However, graft seeding appears to be less practical, as this process has inherent time constraints related to the conditions of EC cultures. Therefore, the most favourable approach is considered to be in-situ endothelialisation of vascular grafts. In vivo circulating EC, EPC,69 CD133+, CD34+, and CD45 progenitor cells70,71 and CD14+ monocytic cells have been shown to differentiate in to EC.72 Interestingly, cells derived

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from bone marrow can differentiate into mature EC, and have been reported to inhibit IH to a greater extent than EPC, which may indicate further developments in this field.73

10.1.8 Important considerations for blood-contacting devices It is essential to consider the surface properties of both biomaterial and blood as important factors in the performance and patency of blood contacting devices. For example, synthetic cardiovascular grafts should mimic the natural mechanical and biological properties of the native tissue. In their development, the primary concern is to match the mechanical properties of strength, viscoelasticity and compliance. However, biomaterials which match these criteria do not necessarily interact optimally with the blood. Hence, the surface properties of the materials are recognised to be key, when considering thrombogenicity.74 Figure 10.1 is a flow chart illustrating the broad categorisation of the strategies of developing optimally functioning blood-contacting devices. Much effort has focused on the modification of surfaces to optimise antithrombogenic surface properties, and two approaches exist in the development of cardiovascular grafts. The first approach involves the design of a permanent vascular replacement, which has a non-adhesive, inert, non-biofouling surface. Physicochemical methods have been applied to achieve this aim, using electrochemical polishing, surface roughening, ordered patterning, plasma treatment, chemical etching, and passive or covalent surface coatings. The second approach aims to functionalise the grafts in such a way that it facilitates (or activates) a cascade of biological events, which can eventually regenerate or replace functioning tissue.75 Biofunctionalisation of surfaces is a popular research theme, which relies on the tools of biology to create biomimetic surfaces incorporating biologically active (or inactive) molecules to generate a specific response(s). Figure 10.2 is a summary of the principal methods in applied surface modification techniques. In this way, surface modification can be directed towards optimising the following: (1) protein adsorption, (2) the generation of thrombin (and its formation, leading to blood coagulation), (3) platelet adhesion (followed by aggregation and activation), and (4) cellular behaviour at the surface of the prosthesis. All strategies are designed to optimise patency-limiting thrombogenic events at the blood-biomaterial interface. For example, vascular graft endothelialisation has been highlighted as the ultimate solution to the problems and complications of thrombogenicity. In recent years, the surface modification of cardiovascular biomaterials has gained considerable interest in terms of its applications in surgical intervention relating to autologous, synthetic and tissue engineered scaffolds for bypass graft materials. In the following sections, we discuss the modification of such materials by physicochemical techniques and biomimetic approaches to improve graft patency and reduce thrombogenicity. This stimulating field is vast, and the

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ß Woodhead Publishing Limited, 2011 10.1 Broad categorisation of strategies for developing blood-contacting devices according to the types of scaffolds and surface modification techniques of mechanically favourable (non-biodegradable) grafts.80,175

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10.2 Examples of various physical, chemical and biofunctionalisation techniques to enhance haemocompatibility. Biofunctionalised surfaces interact with cell surface receptors, i.e. integrins. Whereas physiochemical modification can influence cell-material interactions through charge, topography and attractive/repulsive forces due to hydrophobic and hydrophilic interactions.

emergence of new articles appears on a regular basis. The authors have chosen a relatively small number of studies from the literature in order to outline each example in detail. The reader is encouraged to locate further examples cited elsewhere in the literature.23,75±79 A considerable amount of research effort has been directed toward inhibiting thrombosis and the formation of IH. Various strategies exist to inhibit these processes and prolong graft patency,79 including the modification of grafts with various anti-coagulants (heparin), anti-platelets factors (glycoprotein IIb/IIIa inhibitors) and anti-proliferating agents (rapamycin).80,81 In the following sections, we consider different surface modification techniques that are designed to minimise complications which arise at the blood-biomaterial interface.

10.2

Physicochemical modification

10.2.1 Surface roughness A range of physical techniques have been applied to modify the surface topography of vascular graft materials. Topography on the micron- and nanometre scale is an important physical property, which influences protein adsorption, platelet adhesion, thrombogenicity and cell behaviour.82 The inclusion of pores, pits, and grooves becomes unavoidable at this scale during the manufacturing process of blood-contacting devices. For example, a recent study revealed that

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the surface roughness of ePTFE graft luminal surfaces was significantly higher (147.0 nm) when compared with external surfaces (1.74 nm).83 Plasma proteins such as fibrinogen have been shown to adhere to nanostructures and bind to platelet receptors more efficiently than flat structures.4 Albumin, fibrinogen and fibronectin all interact with a dialysis membrane's surface topography, which plays a crucial role in the adsorption process.16 Such surfaces have been shown to promote fibronectin and vitronectin adsorption and direct a cascade of interactions from the blood and surrounding tissues.84±86 Surface porosity is a crucial factor when considering the topography of vascular graft materials. A recent study looked at the effect of porosity (ranging from 5 to 90 m in diameter) on EC growth. It was found that EC cell growth was enhanced by smaller pores (5±20 m in diameter) and at a lower interpore distances.87 Electrochemical polishing techniques have been used to investigate the thrombogenicity of coronary stents when compared with non-polished (rough) controls. This study used radioactive 125I-labelled fibrinogen and 51Cr-labelled platelets in various animal models.88 The total thrombus weight and fibrinogen adsorption was significantly lower on polished stents; however, platelet adhesion was found to be similar on both stent surfaces. This study highlighted decreased rates of thrombogenicity and reduced IH on smooth polished stents after implantation. More recently, pioneering studies by Dibra and co-workers studied the relationship of surface topography in 200 patients undergoing implantation of sand-blasted versus smooth stent surfaces.89 Sand-blasting created surface features between 0.09 and 0.21 m. The results suggest that rough surfaces may favour EC adhesion, when compared with conventional smooth stent surfaces.89 Earlier studies investigating the thrombogenicity of intravascular catheters were achieved by roughening the surface by controlling the extrusion process, and found that rough surfaces were more thrombogenic than those with a smoother surface profile.90 It was suggested that the effect of roughness on thrombus formation could be due to differences in the adhesion of thrombi to the surface, rather than differences in thrombogenicity per se.91 Acid etching has been used to create micro-roughened substrates with a 10 m surface profile (composed of 1±2 m scale surface pits). The agglomeration of red blood cells (RBCs) and the adhesion of platelets was more extensive on microroughened surfaces after a 10-minute contact time, when compared with controls.92

10.2.2 Ordered patterning Micro- and nanometre-scale fabrication of surfaces can be achieved by using ebeam photolithography, and by casting synthetic polymers onto etched silicon, resulting in reproducible ordered structures. For example, micro-channels were created on PU films with a defined surface geometry (42 m wide and a depth of 32 m). It was found that coating the micro-structures with fibronectin lowered

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the shear stress at the basal plane surface, preventing cell detachment and maximising EC retention.93 Such patterning would favour EC cell adhesion on vascular grafts under continuous flow. PU films have also been used to create micro-patterns that mimic sub-endothelial ECM topography from healthy blood vessels.94 The EC spread more rapidly with morphology that resembles that of the native arteries, compared with EC cultured on untextured controls. Bettinger and co-workers studied rounded micron-scale features, which were shown to align and elongate endothelial cells to facilitate cell±cell interactions and promote tissue formation.95 PU vascular patches have been patterned using laser excitation to create regular spaced fibres (100 m in length) at the luminal surface.96 The textured surfaces were found to stabilise thrombus formation. Cell migration and tissue healing occurred more rapidly on the textured surface, forming a stable pseudo-neointimal layer, when compared with smooth substrates.96 However, no significant histological difference was apparent between textured and smooth surfaces at 3 weeks, except for a higher proportion of red thrombus and a lower healing rate on smooth surfaces. Nonetheless, more studies are necessary to establish the longer-term patency of ordered (and irregular) topographical features on biomaterials, with particular emphasis on EC behaviour. Changing the surface topography on the micron and nanometre scale also leads to localised changes in surface chemistry, as both physicochemical cues are intrinsically linked.97±99 The primary aim of topographical and chemical surface modification is to encourage desirable protein, cellular, and tissue interactions at the blood-biomaterial interface, thus improving patency and performance of the material, since all are known contributory factors that influence thrombogenicity.

10.2.3 UV light and plasma surface modification (PSM) Chemical modification of the surface can be achieved using ultraviolet light (UV) light in a reactive gaseous environment to create new functional groups and moieties. For example, UV light has been used to modify the surface chemistry of ePTFE, Dacron and a novel nanocomposite polymer composed of polyhedral oligomeric silsequioxane (POSS) and poly(carbonate-urea)urethane (PCU). The materials were exposed to an excited xenon lamp at 172 nm in a reducing ammonia atmosphere.100 The irradiation technique resulted in ±N2 and ±O2 surface groups, while retaining the surface morphology, which increased the proliferation of human umbilical vein endothelial cells (HUVEC) after 5 minutes exposure.100 UV modified graft materials show increased proliferation of the cells, and hence show promise for the endothelialisation of vascular grafts. However, earlier studies using similar experimental conditions with ePTFE and Dacron resulted in increased proliferation of SMC, when compared to HUVECs, which are thought to contribute to the development of IH.101 Furthermore,

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chemical patterning can be achieved by UV light and by using a photo mask to create surface microarrays on ePTFE materials. UV irradiation in a reactive gaseous environment resulted in 100 m circular spots (300 m centre-to-centre distance) and was shown to significantly increase the proliferation of HUVECs in the spots resulting in migration, and further population of the surface.102 Therefore, it may be necessary to consider alternative methods that modify either the luminal, or the external surface, or both, in the development of vascular bypass graft materials. A number of studies have shown that PSM is a well established technique to modify the surface properties of, and influence protein and cellular interactions on PTFE, Dacron and PU.103,104 The use of plasma changes the surface functional groups, topography and wettability, and has been effective in improving the haemocompatibility of cardiovascular biomaterials.105±107 For example, low pressure radio frequency (RF) plasma deposition using reactive ammonia gas (±NH3) on ePTFE resulted in a change from hydrophobic to hydrophilic surface properties. The change in surface chemistry is often associated with an increase in the polar component of surface energy. This enables further surface modification, for example using coatings with peptides108 and collagen109 on to the vascular graft surface. Recent studies have investigated the PSM of PTFE and poly(ether)urethane (PEU) with reactive oxygen (O2), argon (Ar), nitrogen (N2) and ammonia (NH3) gas. Different plasma treatments resulted in moderate etching and increased wettability.105 Blood contact with the untreated materials resulted in the activation of platelets, and the clotting cascade. However, plasma treated PTFE and PEU (in all reactive gas species) resulted in a significant reduction in platelet adhesion. It was concluded that, despite the incorporation of the same chemical groups, the pattern of response to blood was not the same for different polymers in vitro.105

10.2.4 Surface passivation and bulk chemical modification Non-fouling surfaces have been used to prevent protein adsorption and platelet adhesion. Much effort has been focused upon the passivation of materials using polymers to achieve a non-adhesive, non-biofouling surface, such as PEG (polyethylene glycol),110±113 hydrogels (containing dextran),114 and PEO (polyethylene oxide).115,116 For example, ePTFE grafts have been coated with polypropylene sulphide (PPS)-PEG and evaluated in arteriovenous models.117 This study included heparinised and non-heparinised graft perfusion, and evaluated for cell adhesion and thrombus formation. No difference was observed in cell adhesion when compared with controls; however, the surface coating significantly decreased thrombus formation when used in conjunction with heparin.117 Dextrans (hydrophilic polysaccharides) show a similar effect to PEG with regard to protein adsorption. Dextrans, PEG and PEO can be further

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chemically modified along the polymer backbone with cell-selective peptides to promote specific cell adhesion.118±120 Spin coating of the luminal surface of ePTFE was recently achieved using a biodegradable elastomer poly(1,8octanediol citrate) (POC). The POC coatings had no effect on graft compliance and delayed thrombosis in vitro, when compared with controls. This study highlighted that POC-ePTFE grafts maintained EC adhesion and proliferation of porcine cells similar to that of the native tissues, and within 10 days the EC were confluent, while only random patches were evident on ePTFE controls.76,121 Another approach to surface modification is modifying the bulk composition to incorporate anti-thrombogenic materials in to the graft. In our group, we have developed a nanocomposite polymer for small diameter bypass graft design, which incorporates POSS. The POSS structures assemble at the surface of PCU to create nanotopographical features, which may be responsible for antithrombogenic surface effects, including the inhibition of factor X activity, effecting blood coagulation, increased clot strength, lysis and reduced platelet adsorption.122,123 More recently, POSS-PCU materials have demonstrated the potential for EC retention and endothelialisation of the graft at the luminal surface, and continue to show promise.124

10.2.5 Bioactive agents A significant number of grafts encounter early graft occlusion, despite the use of anti-coagulants.125 NO produced by EC is found to be effective in preventing thrombosis, the activation of leucocytes (initiating inflammation), limiting platelet aggregation and preventing SMC proliferation,126 improving graft patency. NO releasing vascular implants have become a subject of keen interest.127,128 Various PU grafts have been modified to release NO, preventing platelet adhesion and activation in vitro.129 L-arginine (a NO precursor) has been used to coat vascular grafts and tested in a rabbit model,130 whilst spermine/NO, which was applied on balloon injured arteries of rats131 and a rabbit model132 has been reported to reduce IH. Various NO donors, such as diazeniumdiolates133±135 and s-nitrosothiols,136,137 have been incorporated into biomaterials for controlled release of NO. When diazeniumdiolates (also known as NONOates) are exposed to physiological fluids, they spontaneously release NO (2 molar equivalents) at a known rate. They have a half-life range of between 2 seconds and 17 days, depending on their composition, which may be finely adjusted to suit different applications. S-nitrosothiol has the ability to release NO when in contact with physiological fluids. However, this process does not occur spontaneously.138 They release NO by three known mechanisms: copper ion-mediated decomposition, direct reaction with ascorbate, and homeolytic cleavage by light.138

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10.2.6 Immunosuppressant, cancer chemotherapy and antiplatelet drugs The use of drugs such as Sirolimus (a known immunosuppressant) and Paclitaxel (a mitotic inhibitor used in cancer chemotherapy) has received much attention in blood-contacting devices with the purpose of providing localised antithrombogenic surface properties. Sirolimus and Paclitaxel have both been used for drug-eluting stents and are proven to be effective in the reduction of stenosis. In recent studies, researchers have examined the application of dip-coating both Sirolimus and Paclitaxel drugs on bypass grafts, and both illustrate favourable interactions.76 Paclitaxel coated ePTFE, despite demonstrating differences in haemodynamic behaviour when compared with arterial bypass grafts, showed a mean percentage reduction of 10.4% luminal stenosis, compared with 60.5% on control ePTFE grafts.139 A further study examined coated Sirolimus ePTFE grafts and found a significant reduction in cross-sectional narrowing, compared with untreated ePTFE.80 A recent in vivo study has observed the effect of heparinsirolimus coated onto small diameter PU grafts. The microstructure incorporated collagen and hyaluronic acid in the pores to provide surfaces with cell adhesive and proliferative properties. The heparin-sirolimus grafts stimulated luminal endothelialisation without excessive IH development, with a 100% patency rate after 6 month implantation.140 Photo immobilisation of Persantin (a powerful inhibitor of platelet activation and aggregation) on PU (Pellathane sheets D-55) has been shown to prevent thrombin generation, with long lag times compared with controls, and overall reduced platelet adhesion when applied to small diameter bypass grafts to establish a blood compatible luminal surface.141 This technique may have advantages over biological surface modification (biofunctionalisation) techniques, since it would reduce the host immunological response, although careful attention to the localised and systemic long-term pharmacological effects needs to be considered, particularly in the development of IH.

10.3

Biofunctionalisation

Biofunctionalisation of bypass graft materials is a popular research theme. It involves the creation of surfaces which incorporate or display biologically active (or inactive) (macro)molecules at the blood-biomaterial interface (i.e. membrane proteins, lipids, peptides, growth factors, hormones, cell signalling receptors or ligands) to generate specific biochemical responses to the cells and tissues surrounding the vasculature.

10.3.1 Passive functionalisation ECM peptide sequences have been shown to influence cell behaviour, and have been isolated and grafted onto materials to enhance their biological properties.

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For example, fibronectin sequences (REDV, PHSRN, RGD and GRGDSP),142,143 laminin-derived recognition motifs (IKLLI, IKVAV, LRE, PDSGR, RGD, YIGSR), and collagen type I derived sequences (DGEA, Tenascin-C-derived peptides D5 and D50 ).144,145 These peptide ligands can directly interact with cell surface receptors. Of all of the peptides investigated, the RGD peptide has perhaps featured in the largest number of biomaterials studies. In cardiovascular bioengineering, biofunctionalisation of POSS-PCU with RGD and heparin has been reported to increase EC and EPC adhesion.146±151 In addition to the direct incorporation of peptides in to biomaterials, the engineered peptides can themselves be used as building blocks. Elastin engineered from its constituent polypeptides and cross-linked to exhibit properties of native elastin is one such example.152 The bottom-up approach used by nature to create the biopolymers of the ECM is increasingly being imitated by biomaterials scientists in the fabrication of new synthetic materials and scaffolds for creating vascular grafts.153,154 Thrombomodulin (TM) is a transmembrane glycoprotein expressed on the surface of EC cells. It plays a major role in the anti-coagulant system as a cofactor in thrombin induced activation of protein C, and has received considerable attention in association with vascular grafts. When TM interacts with and binds to thrombin, it has no effect on blood coagulation. TM has been incorporated into lipid vesicles and, using a layer-by-layer (LBL) approach, has been immobilised on polyelectrolyte multilayer (PEM) thin films. TM/PEM films were studied under flowing conditions and were shown to maintain their catalytic activity, providing considerable anti-coagulant activity.155 In a further study, TM and heparin, immobilised on the surface of mimetic films, were exposed to a continuous flow circuit and found to inhibit thrombin activity, showing great potential for improving the thrombogenic profile of bloodcontacting devices.156 Polyelectrolyte multilayer (PEM) films using LBL technology is a novel technique, which enables the release of biochemicals in a controlled manner at the blood-biomaterial interface. Multilayer coatings are assembled using electrostatic interactions to incorporate polypeptides, and biodegradable polymers for the controlled delivery of macromolecules. Controlled release of vascular endothelial growth factor (VEGF) has been accomplished using PEM films.157,158 RepiferminÕ (a recombinant form of fibroblast growth factor) is a drug used in wound healing, which has been shown to increase the proliferation of EC, when compared with RepiferminÕ alone.159 EC adhesion has also been shown to increase with PEM films of poly(allylamine hydrochloride).160 A study has shown that the change in redox potential controls the release of macromolecules from PEM films,161 which may be applied to release anti-coagulation factors in the vicinity of vascular bypass grafts. Furthermore, a LBL approach can be performed in a one-step process using a co-axial multi-needle device to deposit layers by electrohydrodynamic jetting.162 In addition to multilayer

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coatings, this technique creates fibres and tubular scaffolds, which can aid in the design of tissue engineered, bypass grafts coupled with targeted drug delivery. However, the release of growth factors raises concerns about their stability, because of the rapid degradation and high turnover rate in vivo.

10.3.2 Covalently linked macromolecules A recent study investigated the modification of PU materials by incorporating branched polyethylenimine (PEI) into the PU backbone, followed by the covalent attachment of hyaluronic acid (HA), heparin, and PEG (as a nonadhesive control).163 PU materials modified with HA were more effective at preventing protein adsorption and platelet adhesion with confluent EC monolayers on low molecular weight HA, but not high molecular weight HA, heparin or PEG. The use of drug-eluting materials to prevent thrombus formation and IH prior to endothelialisation has led to the use of a variety of anti-thrombogenic and anti-coagulant agents to be incorporated onto the surface of bypass grafts.79,164,165 The Food and Drug Administration (FDA) has approved several commercially available grafts which have used heparin. Heparin is composed of a highly sulphated glycosaminoglycan, and was the first anti-coagulant to be immobilised on surfaces to improve graft patency, as a potent inhibitor of thrombin generation (via anti-thrombin III). A randomised study compared the patency of heparin-modified Dacron and PTFE, and observed 55% vs. 42% patency after three years.166 A 5-year clinical trial with 209 patients has been carried out to test the efficacy of heparin-modified Dacron and PTFE for femoropopliteal bypass* and has shown better patency rates with modified Dacron than PTFE at 3 years, but no difference after 5 years, where the primary patency rate was 46% for modified Dacron, compared with 35% for PTFE.167 A further study of the heparin-modified ePTFE grafts has been performed with 86 patients as femoropopliteal bypass provided promising early patency and limb salvage results.168 But for longer-term use, there are concerns associated with the side effects of heparin related to haemostasis and an increased risk of premature bleeding.169 A naturally occurring thrombin inhibitor, hirudin, is used for therapeutics by isolating it from the salvia of medicinal leeches. Researchers have shown that Dacron, covalently modified with hirudin, had less thrombus formation than Dacron.170,171 Unlike heparin, hirudin is not dependent on heparin cofactors, and anti-heparin proteins do not inactivate hirudin.170,171 Hirudin has a localised anti-coagulation effect, weak allergenicity, and little to no effect on the EC layer.80,170 However, relatively few studies have been carried out into the anti-

*

A femoropopliteal bypass is a prosthesis that bypasses an obstruction in the femoral artery.

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thrombogenic potential of hirudin. One study used albumin as a coating to study the anti-thrombogenic effects of bound hirudin on Dacron. Dacron grafts were exposed to non-heparinised blood for a period of 2 hours while implanted in a canine model. The results showed less thrombus formation and a reduced amount of IH compared with controls.172 In addition, further studies have shown that hirudin bound to carboxylated PU reduced thrombogenicity. However, these reports have been established in vitro, and no in vivo studies have been conducted to date, although they continue to attract interest and show considerable promise.173

10.3.3 Peptide capture ± cell-based strategies Antibody mediated capture techniques for recruiting stem cells is a methodical approach to stem cell technology for applications in vascular bypass graft endothelialisation. In Fig. 10.3, the need for in-situ endothelialisation has led to a considerable interest in stem cells, which have the potential to transform into functioning cells. EC and EPC have been found to express CD34+ and antibodies raised against this epitope can be attached onto bypass graft surfaces to facilitate interactions between the graft and progenitor cells. Anti±human CD34 monoclonal antibodies (IgG2a, epitope class III) were immobilised on ePTFE graft materials using a multi-step process. The antibodies resulted in rapid capture and complete endothelialisation within 72 hours of seeding.174 However, despite the endothelial graft coverage, IH was observed at 4 weeks post implantation, and therefore further studies are necessary to investigate the overall cellular and molecular biological effects of this capture method.174 This technique is currently applied in stent grafts, and clinical trials are investigating the effect of CD34+ capturing bypass grafts.

10.4

Conclusions

In summary, when biomaterials interact with blood, they adsorb proteins which promote interactions with platelet receptors and lead to a cascade of events. The process eventually leads to platelet adhesion, aggregation and thrombus formation, which result in graft failure. The arterial wall presents a natural, well-orchestrated non-thrombogenic environment, which prevents triggering of platelet activation and the clotting cascade. Surface modification methods pave the way to achieve biomimetic, haemocompatible vascular implants, and thus these applications play a significant role in preserving the function of vascular implants. Surface modification can be applied to control vascular interactions between cells and blood-contacting devices. Natural endothelium is considered to be a role model for ideal surface modifications of vascular implants, and therefore there has been a keen interest in surface endothelialisation and particular interest in NO release to induce a surface with anti-thrombogenic

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ß Woodhead Publishing Limited, 2011 10.3 In situ graft endothelialisation. Bioresponsive vascular grafts can target several biological processes to promote in situ endothelialisation, including (1) promoting the mobilisation of EPC from the bone marrow, (2) encouraging cell-specific (circulating EC, EPC, and stem cells) homing to the vascular graft site, (3) providing cell-specific adhesion motifs on the vascular graft (of a predetermined spatial concentration), and (4) directing the behaviour of the cells post adhesion to rapidly form a mature fully functioning endothelium capable of self-repair.176

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potential. It is well understood that optimal surface modification of vascular implants is closely linked with patency rates. The quantity of drugs incorporated into the graft, and how long it should be available in vivo, is a matter of significant importance. In addition, it is essential to understand the effects that could result from the use of drug-releasing grafts. Subsequently, if the drug is covalently attached to the grafts, the systemic effect would be minimised due to the drug being released into circulation, although possible localised side effects also need to be taken into consideration. There is a great clinical significance for vascular implants having optimal biomechanical and surface properties that play a major role in bringing novel bypass grafts for surgical intervention, and provide a valuable link in life saving procedures.

10.5

References

1. Allender S, Peto V, Scarborough P, Kaur A, Rayner M. Coronary heart disease statistics. British Heart Foundation. 2008. 2. Urban P, De Benedetti E. Thrombosis: the last frontier of coronary stenting? Lancet 2007; 369(9562): 619±621. 3. Hannan EL, Racz MJ, Walford G, Jones RH, Ryan TJ, Bennett E Culliford AT, Isom OW, Gold JP, Rose EA. Long-term outcomes of coronary-artery bypass grafting versus stent implantation. N Engl J Med 2005; 352(21): 2174±2183. 4. Hannan EL, Wu C, Walford G, Culliford AT, Gold JP, Smith CR Higgins RS, Carlson RE, Jones RH. Drug-eluting stents vs. coronary-artery bypass grafting in multivessel coronary disease. N Engl J Med 2008; 358(4): 331±341. 5. Seung KB, Park DW, Kim YH, Lee SW, Lee CW, Hong MK, Park SW, Yun SC, Gwon HC, Jeong MH, Jang Y, Kim HS, Kim PJ, Seong IW, Park HS, Ahn T, Chae IH, Tahk SJ, Chung WS, Park SJ. Stents versus coronary-artery bypass grafting for left main coronary artery disease. N Engl J Med 2008; 358(17): 1781±1792. 6. Park DW, Yun SC, Lee SW, Kim YH, Lee CW, Hong MK, Kim JJ, Choo SJ, Song H, Chung CH, Lee JW, Park SW, Park SJ. Long-term mortality after percutaneous coronary intervention with drug-eluting stent implantation versus coronary artery bypass surgery for the treatment of multivessel coronary artery disease. Circulation 2008; 117(16): 2079±2086. 7. Faries PL, Logerfo FW, Arora S, Hook S, Pulling MC, Akbari CM, Campbell DR, Pomposelli FB Jr. et al. A comparative study of alternative conduits for lower extremity revascularization: all-autogenous conduit versus prosthetic grafts. J Vasc Surg 2000; 32(6): 1080±1090. 8. Baguneid MS, Seifalian AM, Salacinski HJ, Murray D, Hamilton G, Walker MG. Tissue engineering of blood vessels. Br J Surg 2006; 93(3): 282±290. 9. Kannan RY, Salacinski HJ, Butler PE, Hamilton G, Seifalian AM. Current status of prosthetic bypass grafts: a review. J Biomed Mater Res B Appl Biomater 2005; 74(1): 570±581. 10. Haruguchi H, Teraoka S. Intimal hyperplasia and hemodynamic factors in arterial bypass and arteriovenous grafts: a review. J Artif Organs 2003; 6(4): 227±235. 11. Sarkar S, Salacinski HJ, Hamilton G, Seifalian AM. The mechanical properties of infrainguinal vascular bypass grafts: their role in influencing patency. Eur J Vasc Endovasc Surg 2006; 31(6): 627±636.

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Surface modification of biomaterials

12. Tiwari A, Cheng KS, Salacinski H, Hamilton G, Seifalian AM. Improving the patency of vascular bypass grafts: the role of suture materials and surgical techniques on reducing anastomotic compliance mismatch. Eur J Vasc Endovasc Surg 2003; 25(4): 287±295. 13. Tengvall P, Ingemar L, Liedberg B. Protein adsorption studies on model organic surfaces: an ellipsometric and infrared spectroscopic approach. Biomaterials 1998; 19: 407±422. 14. Horbett TA, Ratner BD, Schakenraad JM, Schoen FJ. Biomaterials science: an introduction to materials in medicine. Edited by B. D. Ratner, A. S. Hoffman, F. J. Schoen, J. E. Lemons (Academic Press, New, York, 1996), Ch 3, pp. 133±141. 15. Vroman L. Effect of adsorbed proteins on the wettability of hydrophilic and hydrophobic solids. Nature 1962; 196: 476±477. 16. Sutherland DS, Broberg M, Nygren H, Kasemo B. Influence of nanoscale surface topography and chemistry on the functional behaviour of an adsorbed model macromolecule. Macromolecular Bioscience 2001; 1(6): 270±273. 17. Furlan M. Sticky and promiscuous plasma proteins maintain the equilibrium between bleeding and thrombosis. Swiss Med Wkly 2002; 132(15±16): 181±189. 18. Watson SP. Platelet activation by extracellular matrix proteins in haemostasis and thrombosis. Curr Pharm Des 2009; 15(12): 1358±1372. 19. Gorbet MB, Sefton MV. Biomaterial-associated thrombosis: roles of coagulation factors, complement, platelets and leukocytes. Biomaterials 2004; 25: 5681±5703. 20. Vogler EA, Siedlecki CA. Contact activation of blood-plasma coagulation. Biomaterials 2009; 30(10): 1857±1869. 21. Sefton MV, Gemmell CH, Gorbet MB. What really is blood compatibility? J Biomater Sci Polym Ed 2000; 11(11): 1165±1182. 22. Sarkar S, Sales KM, Hamilton G, Seifalian AM. Addressing thrombogenicity in vascular graft construction. J Biomed Mater Res B: Appl Biomater 2007; 82(1): 100±108. 23. Li L, Terry CM, Shiu YT, Cheung AK. Neointimal hyperplasia associated with synthetic haemodialysis grafts. Kidney Int 2008; 74(10): 1247±1261. 24. Tesfamariam B. Platelet function in intravascular device implant-induced intimal injury. Cardiovasc Revasc Med 2008; 9(2): 78±87. 25. Williams DF. The Williams Dictionary of Biomaterials. Liverpool University Press, Liverpool, 1999. 26. Mironov V, Kasyanov V, Markwald RR. Nanotechnology in vascular tissue engineering: from nanoscaffolding towards rapid vessel biofabrication. Trends Biotechnol 2008; 26(6): 338±344. 27. Miller DC, Webster TJ, Haberstroh KM. Technological advances in nanoscale biomaterials: the future of synthetic vascular graft design. Expert Rev Med Devices 2004; 1(2): 259±268. 28. Miller DC, Thapa A, Haberstroh KM, Webster TJ. Endothelial and vascular smooth muscle cell function on poly(lactic-co-glycolic acid) with nano-structured surface features. Biomaterials 2004; 25(1): 53±61. 29. Vane JR, Botting RM. Endothelium ± thromboresistance. Wien Klin Wochenschr 1991; 103(14): 403±404. 30. Miller DC, Haberstroh KM, Webster TJ. Mechanism(s) of increased vascular cell adhesion on nanostructured poly(lactic-co-glycolic acid) films. J Biomed Mater Res A 2005; 73A(4): 476±484. 31. Miller DC, Haberstroh KM, Webster TJ. PLGA nanometer surface features manipulate fibronectin interactions for improved vascular cell adhesion. J Biomed

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with blood

275

Mater Res A 2007; 81A(3): 678±684. 32. Hoffmann J, Groll J, Heuts J, Rong H, Klee D, Ziemer G, Moeller M, Wendel HP. Blood cell and plasma protein repellent properties of star-PEG-modified surfaces. J Biomater Sci Polym Ed 2006; 17(9): 985±996. 33. Ochsenhirt SE, Kokkoli E, McCarthy JB, Tirrell M. Effect of RGD secondary structure and the synergy site PHSRN on cell adhesion, spreading and specific integrin engagement. Biomaterials 2006; 27(20): 3863±3874. 34. Massia SP, Hubbell JA. An Rgd spacing of 440 nm is sufficient for integrin alphaV-beta-3-mediated fibroblast spreading and 140 nm for focal contact and stress fiber formation. J Cell Biol 1991; 114(5): 1089±1100. 35. Koo LY, Irvine DJ, Mayes AM, Lauffenburger DA, Griffith LG. Co-regulation of cell adhesion by nanoscale RGD organization and mechanical stimulus. J Cell Sci 2002; 115(7): 1423±1433. 36. Patel S, Tsang J, Harbers GM, Healy KE, Li S. Regulation of endothelial cell function by GRGDSP peptide grafted on interpenetrating polymers. J Biomed Mater Res A 2007; 83(2): 423±433. 37. Schense JC, Hubbell JA. Three-dimensional migration of neurites is mediated by adhesion site density and affinity. J Biol Chem 2000; 275(10): 6813±6818. 38. Koivunen E, Wang BC, Ruoslahti E. Phage libraries displaying cyclic-peptides with different ring sizes ± ligand specificities of the RGD-directed integrins. BioTechnology 1995; 13(3): 265±270. 39. Pierschbacher MD, Ruoslahti E. Influence of stereochemistry of the sequence ArgGly-Asp-Xaa on binding-specificity in cell-adhesion. J Biol Chem 1987; 262(36): 17294±17298. 40. Xiao Y, Truskey GA. Effect of receptor-ligand affinity on the strength of endothelial cell adhesion. Biophys J 1996; 71(5): 2869±2884. 41. Cheng S, Craig WS, Mullen D, Tschopp JF, Dixon D, Pierschbacher MD. Design and synthesis of novel cyclic RGD-containing peptides as highly potent and selective integrin alpha IIb beta 3 antagonists. J Med Chem 1994; 37(1): 1±8. 42. Hsu SH, Chu WP, Lin YS, Chiang YL, Chen DCH, Tsai CL. The effect of an RGDcontaining fusion protein CBD-RGD in promoting cellular adhesion. J Biotechnol 2004; 111(2): 143±154. 43. Maheshwari G, Brown G, Lauffenburger DA, Wells A, Griffith LG. Cell adhesion and motility depend on nanoscale RGD clustering. J Cell Sci 2000; 113(10): 1677±1686. 44. Hynes RO. Integrins: versatility, modulation, and signaling in cell adhesion. Cell 1992; 69(1): 11±25. 45. Smith JW, Cheresh DA. Integrin (alpha v beta 3)-ligand interaction. Identification of a heterodimeric RGD binding site on the vitronectin receptor. J Biol Chem 1990; 265(4): 2168±2172. 46. Smith JW, Ruggeri ZM, Kunicki TJ, Cheresh DA. Interaction of integrins alpha v beta 3 and glycoprotein IIb-IIIa with fibrinogen. Differential peptide recognition accounts for distinct binding sites. J Biol Chem 1990; 265(21): 12267±12271. 47. Cheng S, Craig WS, Mullen D, Tschopp JF, Dixon D, Pierschbacher MD. Design and synthesis of novel cyclic RGD-containing peptides as highly potent and selective integrin alpha IIb beta 3 antagonists. J Med Chem 1994; 37(1): 1±8. 48. Humphries JD, Byron A, Humphries MJ. Integrin ligands at a glance. J Cell Sci 2006; 119(19): 3901±3903. 49. Ruegg C, Dormond O, Mariotti A. Endothelial cell integrins and COX-2: mediators and therapeutic targets of tumor angiogenesis. Biochim Biophys Acta 2004; 1654(1): 51±67.

ß Woodhead Publishing Limited, 2011

276

Surface modification of biomaterials

50. Urbich C, Dimmeler S. Endothelial progenitor cells: characterization and role in vascular biology. Circ Res 2004; 95(4): 343±353. 51. Chavakis E, Aicher A, Heeschen C, Sasaki K, Kaiser R, El Makhfi N, Urbich C, Peters T, Scharffetter-Kochanek K, Zeiher AM, Chavakis T, Dimmeler S. Role of beta2-integrins for homing and neovascularization capacity of endothelial progenitor cells. J Exp Med 2005; 201(1): 63±72. 52. Rodenberg EJ, Pavalko FM. Peptides derived from fibronectin type III connecting segments promote endothelial cell adhesion but not platelet adhesion: implications in tissue-engineered vascular grafts. Tissue Eng 2007; 13(11): 2653±2666. 53. Ashman LK. The biology of stem cell factor and its receptor C-kit. Int J Biochem Cell Biol 1999; 31(10): 1037±1051. 54. Dentelli P, Rosso A, Balsamo A, Colmenares BS, Zeoli A, Pegoraro M, Camussi G, Pegoraro L, Brizzi MF. C-KIT, by interacting with the membrane-bound ligand, recruits endothelial progenitor cells to inflamed endothelium. Blood 2007; 109(10): 4264±4271. 55. Hristov M, Zernecke A, Bidzhekov K, Liehn EA, Shagdarsuren E, Ludwig A, Weber C. et al. Importance of CXC chemokine receptor 2 in the homing of human peripheral blood endothelial progenitor cells to sites of arterial injury. Circ Res 2007; 100(4): 590±597. 56. Bunting S, Moncada S, Vane JR. Antithrombotic properties of vascular endothelium. Lancet 1977; 2(8047): 1075±1076. 57. Hoenig MR, Campbell GR, Campbell JH. Vascular grafts and the endothelium. Endothelium 2006; 13(6): 385±401. 58. Zilla P, von OU, Deutsch M. The endothelium: a key to the future. J Card Surg 1993; 8(1): 32±60. 59. Seifalian AM, Tiwari A, Rashid ST, Salacinski H, Hamilton G. Impregnation of the the polymeric graft with adhesives molecules, typically oligopeptides or glycoprotein improves retention. Artif Organs 2002; 26(2): 209±210. 60. Meinhart JG, Deutsch M, Fischlein T, Howanietz N, Froschl A, Zilla P. Clinical autologous in vitro endothelialization of 153 infrainguinal ePTFE grafts. Ann Thorac Surg 2001; 71(5): S327±S331. 61. Thomas AC, Campbell GR, Campbell JH. Advances in vascular tissue engineering. Cardiovasc Pathol 2003; 12(5): 271±276. 62. Thebaud NB, Pierron D, Bareille R, Le Visage C, Letourneur D, Bordenave L. Human endothelial progenitor cell attachment to polysaccharide-based hydrogels: a pre-requisite for vascular tissue engineering. J Mater Sci-Mater Med 2007; 18(2): 339±345. 63. Salacinski HJ, Tai NR, Punshon G, Giudiceandrea A, Hamilton G, Seifalian AM. Optimal endothelialisation of a new compliant poly(carbonate-urea)urethane vascular graft with effect of physiological shear stress. Eur J Vasc Endovasc Surg 2000; 20(4): 342±352. 64. Herring M, Gardner A, Glover J. Single-staged technique for seeding vascular grafts with autogenous endothelium. Surgery 1978; 84(4): 498±504. 65. Tiwari A, Salacinski HJ, Hamilton G, Seifalian AM. Tissue engineering of vascular bypass grafts: Role of endothelial cell extraction. Eur J Vasc Endovasc Surg 2001; 21(3): 193±201. 66. Kaushal S, Amiel GE, Guleserian KJ, Shapira OM, Perry T, Sutherland FW, Rabkin E, Moran AM, Schoen FJ, Atala A, Soker S, Bischoff J, Mayer JE Jr. Functional small-diameter neovessels created using endothelial progenitor cells expanded ex vivo. Nat Med 2001; 7(9): 1035±1040.

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with blood

277

67. DiMuzio P, Tulenko T. Tissue engineering applications to vascular bypass graft development: the use of adipose-derived stem cells. J Vasc Surg 2007; 45 Suppl A: A99±103. 68. Bordenave L, Fernandez P, Remy-Zolghadri M, Villars S, Daculsi R, Midy D. In vitro endothelialized ePTFE prostheses: clinical update 20 years after the first realization. Clin Hemorheol Microcirc 2005; 33(3): 227±234. 69. Melero-Martin JM, Khan ZA, Picard A, Wu X, Paruchuri S, Bischoff J. In vivo vasculogenic potential of human blood-derived endothelial progenitor cells. Blood 2007; 109(11): 4761±4768. 70. Urbich C, Dimmeler S. Endothelial progenitor cells: characterization and role in vascular biology. Circ Res 2004; 95(4): 343±353. 71. Jevon M, Dorling A, Hornick PI. Progenitor cells and vascular disease. Cell Proliferation 2008; 41: 146±164. 72. Krenning G, Dankers PYW, Jovanovic D, van Luyn MJA, Harmsen MC. Efficient differentiation of CD14(+) monocytic cells into endothelial cells on degradable biomaterials. Biomaterials 2007; 28(8): 1470±1479. 73. Fujiyama S, Amano K, Uehira K, Yoshida M, Nishiwaki Y, Nozawa Y Jin D, Takai S, Miyazaki M, Egashira K, Imada T, Iwasaka T, Matsubara H. Bone marrow monocyte lineage cells adhere on injured endothelium in a monocyte chemoattractant protein-1-dependent manner and accelerate reendothelialization as endothelial progenitor cells. Circ Res 2003; 93(10): 980±989. 74. Kannan RY, Salacinski HJ, Vara DS, Odlyha M, Seifalian AM. Review paper: principles and applications of surface analytical techniques at the vascular interface. J Biomater Appl 2006; 21(1): 5±32. 75. Sarkar S, Schmitz-Rixen T, Hamilton G, Seifalian AM. Achieving the ideal properties for vascular bypass grafts using a tissue engineered approach: a review. Med Biol Eng Comp 2007; 45(4): 327±336. 76. Kapadia MR, Popowich DA, Kibbe MR. Modified prosthetic vascular conduits. Circulation 2008; 117(14): 1873±1882. 77. Jordan SW, Chaikof EL. Novel thromboresistant materials. J Vasc Surg 2007; 45(6) Sup 1: 104±115. 78. Hoenig MR, Campbell GC, Campbell JH. Vascular grafts and the endothelium. Endothelium 2006, 13(6): 385±401. 79. Sarkar S, Sales KM, Hamilton G, Seifalian AM. Addressing thrombogenicity in vascular graft construction. J Biomed Mater Res B Appl Biomater 2007; 82(1): 100±108. 80. Cagiannos C, Abul-Khoudoud O, Shell DH, DeRijk W, Jennings LK, Handorf CR, Fabian TC. Rapamycin-coated expanded polytetrafluoroethylene bypass grafts exhibit decreased anastomotic neointimal hyperplasia in a porcine model. J Vasc Surg 2005; 42(5): 980±988. 81. Kidane AG, Salacinski H, Tiwari A, Bruckdorfer KR, Seifalian AM. Anticoagulant and antiplatelet agents: their clinical and device application(s) together with usages to engineer surfaces. Biomacromolecules 2004; 5(3): 798±813. 82. Jung YL, Donahue HJ. Cell sensing and response to micro- and nanostructured surfaces produced by chemical and topographic patterning. Tissue Eng. 2007; 13(8): 1879±1891. 83. Doble M, Makadia N, Pavithran S, Kumar RS. Analysis of explanted ePTFE cardiovascular grafts (modified BT shunt). Biomed Mater 2008; 3(3): 1±9. 84. Miller DC, Thapa A, Haberstroh KM, Webster TJ. Endothelial and vascular smooth muscle cell function on poly(lactic-co-glycolic acid) with nano-structured surface features. Biomaterials 2004; 25(1): 53±61.

ß Woodhead Publishing Limited, 2011

278

Surface modification of biomaterials

85. Miller DC, Haberstroh KM, Webster TJ. Mechanism(s) of increased vascular cell adhesion on nanostructured poly(lactic-co-glycolic acid) films. J Biomed Mater Res A 2005; 73A(4): 476±484. 86. Miller DC, Haberstroh KM, Webster TJ. PLGA nanometer surface features manipulate fibronectin interactions for improved vascular cell adhesion. J Biomed Mater Res A 2007; 81A(3): 678±684. 87. Narayan D, Venkatraman SS. Effect of pore size and interpore distance on endothelial cell growth on polymers. J Biomed Mater Res A 2007; 87A(3): 710± 718. 88. De Scheerder I, Sohier J, Wang K, Verbeken E, Zhou XR, Froyen L, Van Humbeeck J, Piessens J, Van de Werf, F. Metallic surface treatment using electrochemical polishing decreases thrombogenicity and neointimal hyperplasia of coronary stents. J Inter Cardiol 2000; 13 (3): 179±185. 89. Dibra A, Kastrati A, Mehilli J, Pache, J., von Oepen, R., Dirschinger, J., SchoÈmig, A. Influence of stent surface topography on the outcomes of patients undergoing coronary stenting: A randomised double-blind controlled trial. Catheterization & Cardiovascular Interventions 2005; 65(3): 374±380. 90. Hecker JF, Scandrett LA. Roughness and thrombogenicity of the outer surfaces of intravascular catheters. J Biomed Mater Res 1985; 19(4): 381±395. 91. Hecker JF, Edwards RO. Effects of roughness on the thrombogenicity of a plastic. J Biomed Mater Res 1981; 15: 1±7. 92. Park JY, Davies JE. Red blood cell and platelet interactions with titanium implant surfaces. Clin Oral Impl Res 2000; 11: 530±539. 93. Daxini SC, Nichol JW, Sieminski AL, Smith G, Gooch KJ, Shastri VP. Micropatterned polymer surfaces improve retention of endothelial cells exposed to flow-induced shear stress. Biorheology 2006; 43(1): 45±55. 94. Goodman SL, Sims PA, Albrecht RM. Three-dimensional extracellular matrix textured biomaterials. Biomaterials 1996; 17: 2087±2095. 95. Bettinger CJ, Orrick B, Misra A, Langer R, Borenstein JT. Microfabrication of poly(glycerol-sebacate) for contact guidance applications. Biomaterials 2006; 27(12): 2558±2565. 96. Fujisawa N, Poole-Warren AL, Woodard JC, Bertram CD, Schindhelm KA. Novel textured surface for blood-contact. Biomaterials 1999; 20: 955±962. 97. Curtis A, Wilkinson C. Nanotechniques and approaches in biotechnology. Trends Biotechnol 2001; 19(3): 97±101. 98. Curtis A, Casey B, Gallagher J, Pasqui D, Wood MA, Wilkinson C. Substratum nanotopography and the adhesion of biological cells. Are symmetry or regularity of nanotopography important? Biophys Chem 2001; 94: 275±283. 99. Curtis A, Wilkinson, C, Crossan J, Broadley C, Darmani H, Johal K, Jorgensen, H, Monaghan W. An In vivo microfabricated scaffold for tendon repair. Eur Cells Mater 2005; 9: 50±57. 100. Olbrich M, Punshon G, Frischauf I, Salacinski HJ, Rebollar E, Romanin C, Seifalian AM, Heitz J. UV surface modification of a new nanocomposite polymer to improve cytocompatibility. J Biomater Sci Polym Ed 2007; 18(4): 453±468. 101. Heitz J, Gumpenberger T, Kahr H, Romanin C. Adhesion and proliferation of human vascular cells on UV-light-modified polymers. Biotechnol Appl Biochem 2004; 39: 59±69. 102. Mikulikova R, Moritz S, Gumpenberger T, Olbrich M, Romanin C, Bacakova L, Svorcik V, Heitz J. Cell microarrays on photochemically modified polytetrafluoroethylene. Biomaterials 2005; 26: 5572±5580.

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with blood

279

103. Khorasani MT, Mirzadeh H, Irani S. Plasma surface modification of poly (L-lactic acid) and poly (lactic-co-glycolic acid) films for improvement of nerve cells adhesion. Rad Phys Chem 2008; 77(3): 280±287. 104. Khorasani MT, MoemenBellah S, Mirzadeh H, Sadatnia B. Effect of surface charge and hydrophobicity of polyurethanes and silicone rubbers on L929 cells response. Colloids and Surfaces B: Biointerfaces 2006; 51(2): 112±119. 105. Rhodes NP, Wilson DJ, Williams RL. The effect of gas plasma modification on platelet and contact phase activation processes. Biomaterials 2007; 28(31): 4561± 4570. 106. Wang C, Chen J, Li, R. Studies on surface modification of poly(tetrafluoroethylene) film by remote and direct Ar plasma. Appl Surf Sci 2008; 254(9): 2882±2888. 107. Wilson DJ, Rhodes NP, Williams RL. Surface modification of a segmented polyetherurethane using low-powered gas plasma and its influence on the activation of the coagulation system. Biomaterials 2003; 24(28): 5069±5081. 108. De S, Sharma R, Trigwell S, Laska B, Ali N, Mazumder MK, Mehta JL. Plasma treatment of polyurethane coating for improving endothelial cell growth and adhesion. J Biomater Sci Polym Ed 16 2005; (8): 973±989. 109. Lim HR, Baek HS, Lee MH, Woo YI, Han D, Han MH, Baik HK, Choi WS, Park KD, Chung K, Park J. Surface modification for enhancing behaviors of vascular endothelial cells onto polyurethane films by microwave-induced argon plasma. Surface & Coatings Technol 2008; 202: 5768±5772. 110. Hoffmann J, Groll J, Heuts J, Rong HT, Klee D, Ziemer G et al. Blood cell and plasma protein repellent properties of Star-PEG-modified surfaces. J Biomater Sci Polym Ed 2006; 17(9): 985±996. 111. Jun HW, West JL. Modification of polyurethaneurea with PEG and YIGSR peptide to enhance endothelialization without platelet adhesion. J Biomed Mater Res B Appl Biomater 2005; 72(1): 131±139. 112. Jun HW, West JL. Endothelialization of microporous YIGSR/PEG-modified polyurethaneurea. Tissue Eng 2005; 11(7±8): 1133±1140. 113. Taite LJ, Yang P, Jun HW, West JL. Nitric oxide-releasing polyurethane-PEG copolymer containing the YIGSR peptide promotes endothelialization with decreased platelet adhesion. J Biomed Mater Res B Appl Biomater 2008; 84(1): 108±116. 114. Letourneur D, Machy D, Pelle A, Marcon-Bachari E, D'Angelo G, Vogel M Chaubet F, Michel JB. Heparin and non-heparin-like dextrans differentially modulate endothelial cell proliferation: in vitro evaluation with soluble and crosslinked polysaccharide matrices. J Biomed Mater Res 2002; 60(1): 94±100. 115. Park K, Shim HS, Dewanjee MK, Eigler NL. In vitro and in vivo studies of PEOgrafted blood-contacting cardiovascular prostheses. J Biomater Sci Polym Ed 2000; 11(11): 1121±1134. 116. Han DK, Park K, Park KD, Ahn KD, Kim YH. In vivo biocompatibility of sulfonated PEO-grafted polyurethanes for polymer heart valve and vascular graft. Artif Organs 2006; 30(12): 955±959. 117. Karrer L, Duwe J, Zisch AH, Khabiri E, Cikirikcioglu M, Napoli A, Goessl A, Schaffner T, Hess OM, Carrel T, Kalangos A, Hubbell JA, Walpoth BH. PPS-PEG surface coating to reduce thrombogenicity of small diameter ePTFE vascular grafts. Int J Artificial Organs 2005; 28(10): 993±1002. 118. Mougin K, Ham AS, Lawrence MB, Fernandez EJ, Hillier AC. Construction of a tethered poly(ethylene glycol) surface gradient for studies of cell adhesion kinetics. Langmuir 2005; 21(11): 4809±4812.

ß Woodhead Publishing Limited, 2011

280

Surface modification of biomaterials

119. Groll J, Fiedler J, Engelhard E, Ameringer T, Tugulu S, Klok HA, Brenner RE, Moeller M. A novel star PEG-derived surface coating for specific cell adhesion. J Biomed Mater Res A 2005; 74A(4): 607±617. 120. Massia SP, Stark J. Immobilized RGD peptides on surface-grafted dextran promote biospecific cell attachment. J Biomed Mater Res 2001; 56(3): 390±399. 121. Yang J, Motlagh D, Allen JB, Webb AR, Kibbe MR, Aalami et al. Modulating expanded polytetrafluoroethylene vascular graft host response via citric acid-based biodegradable elastomers. Adv Mater 2006; 18(12): 1493. 122. Kannan RY, Salacinski HJ, Edirisinghe MJ, Hamilton G, Seifalian AM. Polyhedral oligomeric silsequioxane-polyurethane nanocomposite microvessels for an artificial capillary bed. Biomaterials 2006; 27(26): 4618±4626. 123. Kannan RY, Salacinski HJ, De Groot J, Clatworthy I, Bozec L, Horton M, Butler PE, Seifalian AM. The anti-thrombogenic potential of a polyhedral oligomeric silsesquioxane (POSS) nanocomposite. Biomacromolecules 2006; 7(1): 215±223. 124. Punshon G, Sales KM, Vara DS, Hamilton G, Seifalian AM. Assessment of the potential of progenitor stem cells extracted from human peripheral blood for seeding a novel vascular graft material. Cell Prolif 2008; 41(2): 321±335. 125. Poston RS, Gu J, Brown JM, Gammie JS, White C, Nie L, Pierson RN 3rd, Griffith BP. Endothelial injury and acquired aspirin resistance as promoters of regional thrombin formation and early vein graft failure after coronary artery bypass grafting. J Thorac Cardiovasc Surg 2006; 131(1): 122±130. 126. Radomski MW, Palmer RM, Moncada S. Endogenous nitric oxide inhibits human platelet adhesion to vascular endothelium. Lancet 1987; 2(8567): 1057±1058. 127. Verma S, Marsden PA. Nitric oxide-eluting polyurethane vascular grafts of the future? N Engl J Med 2005; 353(7): 730±731. 128. Varu VN, Tsihlis ND, Kibbe MR. Nitric oxide-releasing prosthetic materials. Vasc Endovasc Surg 2009; 43(2): 121±131. 129. Verma S, Marsden PA. Nitric oxide-eluting polyurethanes ± vascular grafts of the future? N Engl J Med 2005; 353(7): 730±731. 130. Kown MH, Yamaguchi A, Jahncke CL, Miniati D, Murata S, Grunenfelder J, Koransky ML, Rothbard JB, Robbins RC. L-arginine polymers inhibit the development of vein graft neointimal hyperplasia. J Thorac Cardiovasc Surg 2001; 121(5): 971±980. 131. Kaul S, Cercek B, Rengstrom J, Xu XP, Molloy MD, Dimayuga P, Parikh AK, Fishbein MC, Nilsson J, Rajavashisth TB, Shah PK. Polymeric-based perivascular delivery of a nitric oxide donor inhibits intimal thickening after balloon denudation arterial injury: role of nuclear factor-kappaB. J Am Coll Cardiol 2000; 35(2): 493±501. 132. Yin ZL, Dusting GJ. A nitric oxide donor (spermine-NONOate) prevents the formation of neointima in rabbit carotid artery. Clin Exp Pharmacol Physiol 1997; 24(6): 436±438. 133. Jun HW, Taite LJ, West JL. Nitric oxide-producing polyurethanes. Biomacromolecules 2005; 6(2): 838±844. 134. Reynolds MM, Hrabie JA, Oh BK, Politis JK, Citro ML, Keefer LK, Meyerhoff ME. Nitric oxide releasing polyurethanes with covalently linked diazeniumdiolated secondary amines. Biomacromolecules 2006; 7(3): 987±994. 135. Shin JH, Metzger SK, Schoenfisch MH. Synthesis of nitric oxide-releasing silica nanoparticles. J Am Chem Soc 2007; 129(15): 4612±4619. 136. Frost MC, Meyerhoff ME. Controlled photoinitiated release of nitric oxide from polymer films containing S-nitroso-N-acetyl-DL-penicillamine derivatized fumed silica filler. J Am Chem Soc 2004; 126(5): 1348±1349.

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with blood

281

137. Frost MC, Meyerhoff ME. Synthesis, characterization, and controlled nitric oxide release from S-nitrosothiol-derivatized fumed silica polymer filler particles. J Biomed Mater Res A 2005; 72(4): 409±419. 138. Varu VN, Tsihlis ND, Kibbe MR. Basic science review: nitric oxide-releasing prosthetic materials. Vasc Endovasc Surg 2009; 43(2): 121±131. 139. Lee BH, Nam HY, Kwon T, Kim SJ, Kwon GY, Jeon HJ, Lim HJ, Lee WK, Park JS, Ko JY, Kim DJ. Paclitaxel-coated expanded polytetrafluoroethylene haemodialysis grafts inhibit neointimal hyperplasia in porcine model of graft stenosis. Nephrol Dial Transplant 2006; 21(9): 2432±2438. 140. Ishii Y, Sakamoto S, Kronengold RT, Virmani R, Rivera EA, Goldman SM Prechtel EJ, Hill JG, Damiano RJ Jr. A novel bioengineered small-caliber vascular graft incorporating heparin and sirolimus: excellent 6-month patency. J Thorac Cardiovasc Surg 2008; 135(6): 1237±1245. 141. Aldenhoff Y, Blezer R, Lindhout T, Koole LH. Photo-immobilization of dipyridamole (PersantinÕ) at the surface of polyurethane biomaterials: reduction of in vitro thrombogenicity. Biomaterials 1997; 18 (2): 167±172. 142. Koo LY, Irvine DJ, Mayes AM, Lauffenburger DA, Griffith LG. Co-regulation of cell adhesion by nanoscale RGD organization and mechanical stimulus. J Cell Sci 2002; 115(7): 1423±1433. 143. de Mel A, Punshon G, Ramesh B, Sarkar S, Darbyshire A, Hamilton G, Seifalian AM. In situ endothelialization potential of a biofunctionalised nanocomposite biomaterial-based small diameter bypass graft. Biomed Mater Eng 2009; 19(4±5): 317±331. 144. Schense JC, Hubbell JA. Three-dimensional migration of neurites is mediated by adhesion site density and affinity. J Biol Chem 2000; 275(10): 6813±6818. 145. Koivunen E, Wang BC, Ruoslahti E. Phage libraries displaying cyclic-peptides with different ring sizes ± ligand specificities of the Rgd-directed integrins. BioTechnology 1995; 13(3): 265±270. 146. Pierschbacher MD, Ruoslahti E. Influence of stereochemistry of the sequence ArgGly-Asp-Xaa on binding-specificity in cell-adhesion. J Biol Chem 1987; 262(36): 17294±17298. 147. Xiao Y, Truskey GA. Effect of receptor-ligand affinity on the strength of endothelial cell adhesion. Biophys J 1996; 71(5): 2869±2884. 148. Cheng S, Craig WS, Mullen D, Tschopp JF, Dixon D, Pierschbacher MD. Design and synthesis of novel cyclic RGD-containing peptides as highly potent and selective integrin alpha IIb beta 3 antagonists. J Med Chem 1994; 37(1): 1±8. 149. Hsu SH, Chu WP, Lin YS, Chiang YL, Chen DCH, Tsai CL. The effect of an RGDcontaining fusion protein CBD-RGD in promoting cellular adhesion. J Biotechnol 2004; 111(2): 143±154. 150. Maheshwari G, Brown G, Lauffenburger DA, Wells A, Griffith LG. Cell adhesion and motility depend on nanoscale RGD clustering. J Cell Sci 2000; 113(10): 1677± 1686. 151. Hynes RO. Integrins: versatility, modulation, and signaling in cell adhesion. Cell 1992; 69(1): 11±25. 152. Langer R, Tirrell DA. Designing materials for biology and medicine. Nature 2004; 428(6982): 487±492. 153. Kotch FW, Raines RT. Self-assembly of synthetic collagen triple helices. Proc Natl Acad Sciences USA 2006; 103(9): 3028±3033. 154. Stegemann JP, Kaszuba SN, Rowe SL. Review: advances in vascular tissue. Tissue Eng 2007; 13(11): 2601±2613.

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

155. Tseng PY, Jordan SW, Sun XL, Chaikof EL. Catalytic efficiency of a thrombomodulin-functionalized membrane-mimetic film in a flow model. Biomaterials 2006; 27(13): 2768±2775. 156. Tseng PY, Rele SS, Sun XL, Chaikof EL. Membrane-mimetic films containing thrombomodulin and heparin inhibit tissue factor-induced thrombin generation in a flow model. Biomaterials 2006; 27(12): 2637±2650. 157. Matsusaki M, Sakaguchi H, Serizawa T, Akashi M. Controlled release of vascular endothelial growth factor from alginate hydrogels nano-coated with polyelectrolyte multilayer films. J Biomater Sci Polym Ed 2007; 18(6): 775±783. 158. Huang M, Vitharana SN, Peek LJ, Coop T, Berkland C. Polyelectrolyte complexes stabilize and controllably release vascular endothelial growth factor. Biomacromolecules 2007; 8(5): 1607±1614. 159. Huang M, Berkland C. Controlled release of RepiferminÕ from polyelectrolyte complexes stimulates endothelial cell proliferation. J Pharm Sci 2009; 98(1): 268± 280. 160. Boura C, Muller S, Vautier D, Dumas D, Schaaf P, Voegel JC, Stoltz JF, Menu P. Endothelial cell-interactions with polyelectrolyte multilayer films. Biomaterials 2005; 26(22): 4568±4575. 161. Zhong Y, Whittington CF, Haynie DT. Stimulated release of small molecules from polyelectrolyte multilayer nanocoatings. Chem Comm 2007; 14: 1415±1417. 162. Ahmad Z, Zhang HB, Farook U, Edirisinghe M, Stride E, Colombo P. Generation of multilayered structures for biomedical applications using a novel tri-needle coaxial device and electrohydrodynamic flow. J R Soc Interface 2008; 5(27): 1255± 1261. 163. Chuang T, Masters KS. Regulation of polyurethane haemocompatibility and endothelialisation by tethered hyaluronic acid oligosaccharides. Biomaterials 2009; 30: 5341±5351. 164. Kidane AG, Salacinski H, Tiwari A, Bruckdorfer KR, Seifalian AM. Anti-coagulant and anti-platelet agents: their clinical and device application(s) together with usages to engineer surfaces. Biomacromolecules 2004; 5(3): 798±813. 165. Hasegawa T, Okada K, Takano Y, Hiraishi Y, Okita Y. Autologous fibrin-coated small-caliber vascular prostheses improve antithrombogenicity by reducing immunologic response. J Thorac Cardiovasc Surg 2007; 133(5): 1268±76, 1276. 166. D e v i n e C , H o n s B , M c C o l l u m C . H e p a r i n - b o n d e d D a c r o n o r polytetrafluoroethylene for femoropopliteal bypass grafting: a multicenter trial. J Vasc Surg 2001; 33(3): 533±539. 167. Devine C, McCollum C. Heparin-bonded Dacron or polytetrafluorethylene for femoropopliteal bypass: five-year results of a prospective randomized multicenter clinical trial. J Vasc Surg 2004; 40(5): 924±931. 168. Bosiers M, Deloose K, Verbist J, Schroe H, Lauwers G, Lansink W et al. Heparinbonded expanded polytetrafluoroethylene vascular graft for femoropopliteal and femorocrural bypass grafting: 1-year results. J Vasc Surg 2006; 43(2): 313±318. 169. Johnson WC, Williford WO, Corson JD, Padberg FT, Jr. Hemorrhagic complications during long-term postoperative warfarin administration in patients undergoing lower extremity arterial bypass surgery. Vascular 2004; 12(6): 362± 368. 170. Kidane AG, Salacinski H, Tiwari A, Bruckdorfer KR, Seifalian AM. Anti-coagulant and antiplatelet agents: their clinical and device application(s) together with usages to engineer surfaces. Biomacromolecules 2004; 5(3): 798±813. 171. Sarkar S, Sales KM, Hamilton G, Seifalian AM. Addressing thrombogenicity in

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172. 173.

174.

175. 176.

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vascular graft construction. J Biomed Mater Res B Appl Biomater 2007; 82(1): 100± 108. Wyers MC, Phaneuf MD, Rzucidlo EM, Contreras MA, LoGerfo FW, Quist WC. In vivo assessment of a novel dacron surface with covalently bound recombinant hirudin. Cardiovasc Pathol 1999; 8(3): 153±159. Phaneuf MD, Szycher M, Berceli SA, Dempsey DJ, Quist WC, LoGerfo FW. Covalent linkage of recombinant hirudin to a novel ionic poly(carbonate) urethane polymer with protein binding sites: determination of surface antithrombin activity. Artif Organs 1998; 22(8): 657±665. Rotmans JI, Heyligers JM, Verhagen HJ, Velema E, Nagtegaal MM, de Kleijn DP, de Groot FG, Stroes ES, Pasterkamp G. In vivo cell seeding with anti-CD34 antibodies successfully accelerates endothelialization but stimulates intimal hyperplasia in porcine arteriovenous expanded polytetrafluoroethylene grafts. Circulation 2005; 112(1): 12±18. Isenberg BC, Williams C, Tranquillo RT. Small-diameter artificial arteries engineered in vitro. Circ Res 2006; 98(1): 25±35. de Mel A, Jell G, Stevens M, Seifalian AM. Biofunctionalisation of biomaterials for accelerated in-situ endothelialisation: a review. Biomacromolecules 2008; 9(11): 2969±2979.

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Modifying biomaterial surfaces with bioactives to control infection H . J . G R I E S S E R , K . V A S I L E V , H . Y S and S . A . A L - B A T A I N E H , University of South Australia, Australia

Abstract: This chapter discusses representative examples of strategies used for reducing infections associated with biomedical implants and devices, which are caused by the attachment of bacteria and their subsequent biofilm formation on device surfaces. A number of approaches have been shown to produce surfaces/coatings that confer resistance to such bacterial colonisation. Particular emphasis is placed on plasma-based technologies, as plasma polymerisation, in particular, is a technology that is transferable between a large range of different materials and hence can be applied to a variety of biomedical devices and implants without having to re-optimise each time, and without the need for specific substrates. A popular strategy continues to be the use of silver ions or silver nanoparticles in polymer matrices. Another popular approach comprises the use of quaternary amine compounds on surfaces. In both cases, antibacterial activity is indeed obtained, but damage to human cells and tissue remains a concern. Alternative approaches utilising the covalent immobilisation of antibiotic molecules are discussed. Key words: antibacterial surfaces, antimicrobial, antibiotic, silver, quaternary amine.

11.1

Introduction

With increasing life spans in modern societies, the demand for biomedical devices continues to increase, and technology enables the design and development of better or longer-lasting biomedical devices. Numerous products are commercially available. By far the largest market segment of biomedical products sold comprises contact lenses, which are often considered to be relatively simple products, yet for many years suffered from excessive fouling that limited on-eye wear times. Interesting surface science challenges needed solving before modern extended-wear silicone hydrogel lenses achieved market success [1]. It is particularly notable that the first two successful extended wear contact lenses, from Ciba Vision [1] and Bausch & Lomb, both utilised plasma surface engineering methodologies. In this chapter, we will discuss how surface engineering technologies can be utilised to fabricate surfaces and coatings to address the persistent problem of

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bacterial infection of biomedical devices and implants. Particular emphasis will be given to plasma technologies, because plasma polymerisation, in particular, is a technology that is transferable between a large range of different materials and hence can be applied to a variety of biomedical devices and implants without having to re-optimise each time, and without the need for specific substrates such as gold for SAMs that enable model studies but have limited applicability for biomedical devices. Thus, many of the examples given in this chapter include the use of plasma technologies, but obviously analogous chemical and surface science principles apply to the surface engineering of devices and implants using technologies other than plasma. Plasma surface treatments and plasma polymer coatings have achieved success per se in a number of biomaterials surface engineering areas, for example for tissue culture ware (see Chapter 1) and as coatings for extended-wear contact lenses [1]. They have also been used with considerable success for the covalent immobilisation of bioactive molecules intended to elicit specific biological responses (see e.g. [2]). Much of such work aimed either to achieve better cell attachment or tissue integration, or resistance to the adhesion of biological molecules, cells, and tissue. The avoidance of cellular adhesion by fouling resistant layers can of course, also serve to reduce the adhesion of bacteria, and in some applications this may be the most relevant approach. However, in other instances, rather than `passive' anti-adhesive strategies, technologies are required that lead to `active' antibacterial surfaces that deter the adhesion of bacteria and the consequent development of biofilm and implant infection while enabling good interfacing with cells and tissue. Why is the literature on antibacterial surfaces and coatings increasing rapidly? With increasing usage of biomedical devices and implants, the issue of bacterial infections at implants and devices has become manifest and poses limitations on the success of medical interventions and implant integration. Socalled device-related infections (DRIs) arise when bacteria attach and proliferate on surfaces of biomedical devices and implants; they are a considerable problem in implant surgery and also with short-term biomedical devices [3±5]. DRIs pose a substantial health risk to patients, often requiring re-operation and replacement of the infected device, and incur considerable costs to the health care system. While infections on contact lenses and catheters become manifest soon after infection, in the case of implants DRIs often are not detected at an early stage. The severity of DRIs varies greatly among devices [5] and patients but can be serious, or even fatal at times. For example, contact lens related infections are rare in the hygienic environment of modern societies, and most users have the common sense to remove contact lenses upon eye soreness, typical for ocular bacterial infection. On the other hand, there are no such early warning signs with bacterial infection of many implants and devices; the outbreak of an infection is masked by ongoing soreness of tissue after surgery. DRI diagnosis on orthopaedic implants usually occurs when a full-blown infection has already

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caused damage to tissue and host organism. Re-operation of infected implants has at times led to the death of elderly patients weakened by the previous operation or other conditions. Health care systems attempt to minimise the infection risk on short-term biomedical devices by prophylactic measures; for example, catheters are replaced at frequent intervals. Such preventative replacement schedules impose considerable costs to health care systems. With implants, the problem cannot be addressed in the same manner. Improved procedures in sterilization and operating theatres have reduced significantly the frequency of early-stage infections of implants; however, infections that may occur many weeks or months after surgery have barely decreased and continue to pose a serious problem. Such late-stage infections may be caused not by the act of surgery but by planktonic bacteria circulating in the vascular system and eventually reaching an incompletely healed wound site, where they may adhere to the implant surface, multiply, and form a biofilm that eventually leads to infection. Bacterial biofilm colonies produce an exocellular polysaccharide matrix that protects them against antibiotics and the host body's innate defense system [4, 6]. As a result, bacterial biofilms on biomedical device surfaces are much more difficult to eradicate by antibiotics than circulating bacteria; hence the need for surgical removal of a substantial proportion of infected implants. A promising strategy for reducing the occurrence of DRIs is to combat the initial attachment of bacteria to implant and device surfaces. Accordingly, research efforts have focused on the development of antibacterial surfaces and thin coatings that can be applied to biomedical devices to confer resistance to bacterial colonisation. Such antibacterial surfaces/coatings should reduce DRIs while not affecting other properties such as the visual clarity of contact lenses or the flexibility of vascular grafts. Nor should, of course, such coatings have adverse effects on host body fluids, cells, and tissue. In this chapter we discuss strategies for the development of antibacterial surfaces and coatings for biomedical devices. We stress that given the variety of devices and implants as well as causative bacteria, there may not exist a single successful approach; antibacterial strategies may need to be tailored to specific product needs. For some products, such as contact lenses, resistance to the attachment of bacteria is required along with biofouling resistance, whereas for hip and knee implants one wishes to deter bacteria while encouraging the close apposition of host cells for close integration of the implant with human tissue. Such discussion is beyond the scope of this chapter; we will focus on the plasma materials science of the design and fabrication of antibacterial surfaces and coatings. Antibacterial approaches are, of course, also applicable to other fields where bacterial attachment and subsequent biofilm formation cause problems, for example in water filtration and purification systems. There is a substantial body of literature on the reduction or prevention of biofouling in such applications,

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but the most sophisticated strategies are found in the biomaterials literature. Reasons for this are that the consequences of infection arising at implant or device surfaces are much more serious in human health care, and that the biomedical devices/implants market is much less price-sensitive.

11.2

Plasma-based strategies for combating devicerelated infections

Plasmas can be used in a number of ways to support the fabrication of antibacterial surfaces and coatings. In this section we outline various approaches. Some will be discussed fairly briefly while approaches that are more frequently used and more interesting and/or more promising will be discussed in more detail in subsequent sections. Plasma surface treatments to introduce chemical groups that deter bacterial attachment is not a promising strategy as the range of chemical groups available via plasma surface treatments is limited and antibacterial action is more complex than just being affected by a single small chemical group. By analogy with wellknown cationic germicides, ammonia plasma treatment could be envisaged as the best-suited plasma surface treatment, but it is evident from the literature, such as for amine plasma polymers, that amine groups by themselves have rather limited antibacterial effectiveness. In addition, for many polymers the issue of surface rearrangement (hydrophobic recovery) after plasma treatments may limit shelf life [7, 8]. Plasma polymers by themselves have shown rather limited effectiveness for deterring bacterial attachment. In one (unpublished) study, n-heptylamine plasma polymer coatings were applied onto silicone hydrogel extended-wear contact lenses and such lenses tested for adhesion of bacteria associated with eye infections. The observed reductions, of less than 50%, were considered insufficient. To improve antibacterial activity, Jampala et al. [9] alkylated amine groups on an ethylene diamine plasma polymer with hexyl bromide, thus forming quaternary amine cationic structures akin to those found in many germicides. While longer alkyl chains are more effective, they also make the surface much more hydrophobic. For applications such as contact lenses, where a fairly hydrophilic surface is required, quaternary alkyl surfaces from plasma polymers are not suitable. In addition, as discussed in Section 11.4.1, quaternary amine compounds cause adverse effects on human cells. Thus, amine plasma polymers derivatised by alkyl quaternisation appear unsuitable for human medical use, but they may be suitable for other applications such as the combating of membrane fouling. Plasma polymers have, however, attracted interest as carrier matrices for nanoparticles and metal ions. The plasma polymer acts as a reservoir for the outdiffusion of ions that combat bacterial attachment. Nanoparticles and metal ions have also been incorporated into many other release matrices, but key

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advantages of plasma polymers for this purpose are the generally excellent adhesion to substrates/devices and the fact that plasma polymer coatings can be applied at such low thickness that the bulk mechanical properties of a device, such as flexibility, are not altered. In addition, results show that plasma polymers may offer a wider range of tailorable delivery rates. Such work will be discussed in Section 11.3. Analogously, plasma polymers can in principle be used for the storage and release of organic antibiotic molecules. However, while metallic nanoparticles can be added during plasma polymer deposition by concurrent sputtering off a metal target, organic molecules must be added into the plasma polymer layer after deposition, by soaking the plasma polymer coated device in a solution of the organic molecule. Owing to the larger size of organic molecules compared with metal ions, diffusion in and out of plasma polymers is expected to be substantially slower. In addition, solubility may be limited. Thus, a more promising approach is the spreading of nanocrystals of organic antibiotics on a substrate/device surface followed by application of a plasma polymer layer that ensures that the nanocrystals are held in place and that release is not too fast upon contact with biological media. Two studies have utilised plasma polymers for covering nanocrystals and controlling diffusive release [10, 11] while others have used plasma polymer layers as barriers to control outdiffusion of antibiotics from a bulk polymer [12]. Compared to systemic drug delivery, a key advantage of local delivery of antibiotics at a specific site is that high local doses can be administered without exceeding the systemic toxicity level of the drug and risking renal and liver complications, for example. Thus, via coatings, increased doses can be applied at the site of a medical implant. An important issue is the kinetics of release. Fast release provides relatively high doses but short-term action, whereas slow release may not reach the required therapeutic level, and might also lead to bacterial resistance at the release site due to survival of strains that adapt. The released antibiotic must act before a protective exocellular matrix layer protects the growing bacterial colony. Bacteria protected by a biofilm can require a thousand times the antibiotic dose necessary to combat bacteria in suspension. An ideal release coating should provide fast initial release in the first six hours after surgery to protect the site while the immune system is weakened, and continuous `prophylactic' slow release. A release approach may be suitable for short-term protection but is unlikely to be effective in preventing delayed infections that arise after months, for example with orthopaedic implants. Particularly when using plasma polymers, the very limited reservoir space available, compared with the release of antibiotics from bulk polymers or from much thicker, solvent-applied coatings, will limit the duration of effective antibacterial action. Whether by release of silver ions or organic antibiotics, plasma polymer coatings therefore are suited to applications where short-term protection against bacterial colonisation is sought.

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For long-term deterrence of bacterial colonisation of devices, surface chemical compositions must be fabricated in a way that can maintain their function over extended periods of time; for example for orthopaedic implants for up to two years given the very slow healing of wound sites in elderly patients. Permanent attachment resistant surfaces are not exclusive to the need for antibacterial materials; in fact, the biomaterials literature contains a large body of reports on surfaces and coatings that can resist the adsorption of proteins and cells, usually by the grafting of hydrogel polymer layers, with poly(ethylene glycol) (PEG) by far the most prominent. A number of studies have used plasma polymer coatings as interlayers for the covalent grafting of fouling-resistant PEG hydrogel layers [for example, 13±15]. It must be noted, however, that in vivo fouling is more complex than what in vitro models can mimic at present, and even in in vitro tests protein and cell resistance may not directly translate to resistance to bacterial attachment [16]. While, in general, adsorbed biomolecular fouling layers promote bacterial colonisation on surfaces of synthetic materials, some bacteria can attach and proliferate on synthetic surfaces without adsorbed biomolecule layers. Thus, resistance to protein adsorption may not translate to resistance to bacteria. Several hydrogel coatings, for example from PEG [17], polyacrylamide [18], and polylysine-PEG copolymers [19] have shown resistance to bacterial colonisation. Few studies exist that address the question of longer-term stability and performance of protective antibacterial layers. Instructive is a study by Kingshott et al. which showed that physisorbed PEO polymers did not provide lasting reduction in bacterial adhesion, whereas PEO chains covalently attached to a bulk material were stably effective [20]. A likely explanation is that bacteria can act as mega-surfactants with high interfacial affinity for the material surface, displacing physisorbed polymer chains from the bulk material surface, whereas surface-grafted polymer chains resist such displacement. It would therefore seem highly advisable to use the pathway of covalent attachment of antibacterial polymers or coatings onto devices in order to prevent possible loss in biological milieus. For low molecular weight antibacterial compounds, covalent attachment is essential as they are more readily detached from a surface than adsorbed polymers, but the main point to note here is that even polymer molecules may not remain on the device surface when in contact with biological media, in which many surface-active proteins and cellular entities may have sufficient affinity for the surface of biomedical devices to displace physisorbed molecules. Thus, we believe that covalent grafting of suitable thin layers onto implants and medical devices is the only viable option for long-term protection against bacterial infections. Plasma treated surfaces and plasma polymers are of great value as platform interlayers for the covalent grafting of antibacterial small molecules and plasma polymers offer the prospect of ready transfer of the coating/grafting methodology to various substrates/devices. The covalent

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attachment (`grafting') of a molecular layer of antibacterial molecules, discussed in Section 11.4, is advantageous also when seeking regulatory approval for new devices; if it can be ascertained that the antibiotics are durably grafted such as to remain on the device surface, one can eliminate concerns about possible adverse effects due to accumulation of antibiotics in body tissues such as brain, liver and spleen.

11.3

Plasma polymers with incorporated metal nanoparticles or ions

11.3.1 Silver Considerable attention has focused for several decades on the use of silver as an antibacterial agent. The antibacterial properties of silver have been known since antiquity. In the light of increasing incidence of resistance to organic antibiotics, metallic silver, silver compounds and silver nanoparticles have enjoyed increased attention in recent years [21]. A number of commercial products have emerged including wound dressings, bandages for burns and chronic wounds, catheters, and other medical devices that contain or are coated with silver, silver salts, or silver nanoparticles. The efficiency of some of the coatings has, however, been subject to controversy, such as the issue of possible adverse effects on host cells and tissue. A number of factors, both from the coating and the environment, appear to determine the effectiveness of silver-based coating. It was found that coatings of solid metallic silver are least efficient. This is because silver needs to be in its oxidised ionic form (Ag‡) in order to exhibit antibacterial action, and surface oxidation provides a limited supply. Another limitation is that after out-diffusion into biological media, Ag‡ ions can be complexed with chloride ions (Clÿ) to produce AgCl, a white precipitate, which is detrimental for antibacterial efficiency. Another aspect is that in many cases most of the loaded silver is released very quickly from the coating, thus limiting the time of protection but, on the other hand, providing fast burst protection. There have been many studies on various thin polymer films loaded with silver ions or silver nanoparticles designed to produce antibacterial coatings. Silver nanoparticles appear advantageous as they give slower release kinetics overall, and silver ions show a strong initial burst release. For biomedical devices it may be advantageous to use thin polymer films containing silver nanoparticles fully encapsulated within the polymer matrix, as opposed to the situation where larger silver particles partly protrude from the polymer film layer. This can be achieved in several ways. Often silver particles are mixed into a polymer/solvent mix and the mixture is then coated and dried/cured. The advantages of plasma polymerisation for creating the polymeric matrix for nanosilver are that vacuum deposition often creates better interfacial adhesion and better coating uniformity at lower, controllable thicknesses, compared with

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solvent coating processes. However, the issue of limited loading and thus limited duration of antibacterial action needs to be considered. There are several ways to incorporate silver nanoparticles into plasma polymers. One approach is the simultaneous deposition of silver and a plasma polymer matrix, involving the use of plasma polymerisation to form a polymeric thin film while adding silver by sputtering from a silver target; this has been done with PEO-like [22, 23] and siloxane [24] plasma polymers. An alternative approach was recently reported by Vasilev et al. [25] in which plasma polymerisation of n-heptylamine or allylamine was used to create aminerich coatings into which silver ions were in-diffused by placing samples in a solution of AgNO3; the time of immersion controls the loaded amount. The silver ions were then reduced to silver using a reducing agent and the silver atoms coalesced into nanoparticles, detected by a characteristic plasmon band in the UV/Vis absorption spectrum. When placing such samples into an aqueous solution, oxidation causes slow release of silver ions. The rate of release was shown to be controllable by the application of a second plasma polymer layer on top of the layer containing silver nanoparticles, with the thickness of this overlayer providing a convenient means of selecting a desired release rate. In this way, the loading and the release kinetics were shown to be adjustable independently over considerable ranges (Fig. 11.1). An antibacterial coating can also be produced by plasma polymerisation of a phosphine-stabilised silver maleimide complex [26]. This method is attractive in terms of comprising a single step but requires synthesis of the precursor.

11.1 Release of silver monitored by the decrease in the intensity of the Plasmon absorbance band, from silver nanopaticles in n-heptylamine plasma polymer (n-ha pp) with no overlayer (squares), a 6 nm thick n-ha pp overlayer (circles) and a 12 nm thick n-ha pp overlayer (triangles). From [25].

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Other plasma-related technologies have also been used for the fabrication of silver-containing polymer films. Polytetrafluoroethylene can be sputter-deposited with concurrent incorporation of nanosilver [27]. Gray et al. [28] used a remote air plasma to activate polyurethane catheter surfaces for electroless silver coating. While there is still some confusion, the scientific consensus (reviewed recently in [29]) appears to tend towards the view that metallic silver has little activity (despite the mention of contact activation) and that it is silver ions, released from silver metal coatings or polymer materials or coatings doped with silver nanoparticles, that exercise biological functions. These ions may interact with bacterial cell walls, complex with proteins and DNA, and affect their biological functions. Silver appears to be the broadest spectrum antibiotic available, and it does not appear to induce resistance. Different scenarios have been proposed for the mode of action of silver against bacteria; one is that silver ions bind to the bacterial cell membrane and damage it by interfering with membrane receptors and bacterial electron transport (impeding the production of adenosine triphosphate, the cell's energy source). Another scenario is that silver ions bind to bacterial DNA and thus damage cell replication, and cause intracellular formation of insoluble compounds with nucleotides and proteins; particularly the amino acids histidine and cysteine will readily complex with Ag‡ ions. Given the similarity of many structural elements that constitute human and bacterial cells, the question arises whether silver ions also interfere with human cell and tissue functions in comparable way. Indeed, Schrand et al. [30] reported that induction of reactive oxygen species, degradation of mitochondrial membrane integrity, disruption of the actin cytoskeleton, and reduction in proliferation after stimulation with nerve growth factor were observed upon exposing neuroblastoma cells to silver nanoparticles. Equally concerning is that in a recent study it was shown that a thin coating of an antimicrobial composite of poly(4-vinylpyridine)-co-poly(vinyl-N-hexylpyridinium bromide) and AgBr nanoparticles on Tygon elastomer tubes caused disruption and activation of blood platelets, making this coating unsuitable for blood-contacting devices applications [31]. It was, however, not clear whether it is the cationic copolymer or the Ag‡ ions, or both, that caused the adverse effects. In addition, another study found that the extent of adverse effects from silver ions differed for different cell lines [32], which may help explain the inconsistencies in the literature where studies using only one cell line each, and different lines in different studies, have reported conflicting findings. Yet, despite such concerns regarding toxicity to human cells and tissue, and evidence of absence of antibacterial effectiveness in vivo [33, 34], a number of companies are building their business on silver-releasing coatings or nanosilvercontaining products, and research continues on variations of the theme. Perhaps more thought needs to be given to where and when such coatings may be suitable or unsuitable; for example, does it matter if a layer of cells next to an

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implant is killed by antibacterial silver? For some devices, it may, whereas for others, it may not matter. In cases such as antibacterial skin wound dressings the short-term uptake of silver ions appears acceptable given the need to keep skin wounds free of infections. In summary, silver ions are clearly effective against various bacteria in vitro. However, questions regarding their clinical effectiveness and damage to human cells are subject of continuing controversy. Silver-releasing coatings may find some applications, but given the concern of potential side effects from silver forming complexes with proteins and altering their functions, we are of the opinion that application of silver-releasing coatings for biomedical devices must be approached with considerable caution except where much of the silver is likely to be removed from the body, such as with urinary catheters, or where the benefits outweigh the risk, such as for skin wound dressings. Silver should, however, not be used lightly, and it is concerning to consider the possibility of substantial amounts of silver released into the environment from antibacterial socks, washing machine additives, and other applications where chemically less risky alternatives are available, given that transition metal ions can accumulate readily in living organisms such as fish and amphibians.

11.3.2 Other metallic ions Very similar studies as discussed above with silver have also been undertaken with other transition metals/metal ions. A number of transition metal ions are known to possess antibacterial activity, and their release from polymeric coatings can analogously be used to achieve short-term prevention of bacterial attachment to materials. For example, analogous to the incorporation of silver nanoparticles in plasma polymer coatings, copper has been incorporated by concurrent plasma polymerisation and sputtering, using hexamethyldisiloxane as the plasma process monomer [35]. However, relatively high loadings of Cu were found necessary for effective antibacterial action. An alternative approach was reported by Zhang et al. [36] who used plasma immersion ion implantation for the incorporation of Cu ions into polyethylene. Subsequent nitrogen plasma immersion ion implantation was used to crosslink polymer chains and thereby regulate the release rate. Plasma enhanced chemical vapour deposition of organometallic compounds can also be used to fabricate metal ion containing films, for example with cobalt [37]. Clearly, a number of transition metal ions can be released from polymeric matrices to deliver antibacterial action. As for silver ions, however, cytotoxic and systemic toxicity effects need to be considered, as such ions can interfere with a variety of human metabolic processes. For non-biomedical applications, it also needs to be considered whether substantial amounts of transition metal ions can responsibly be discharged into the environment, particularly waterways. In our opinion the release of antibacterial metal ions should only

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be used for limited and well-regulated applications where no reasonable alternatives exist.

11.4

Covalent immobilisation of antibacterial molecules

Covalent grafting of antibacterial compounds onto polymer surfaces has been the subject of considerable research in the search for antibacterial surfaces with longer-lasting effectiveness than is possible via release approaches. The polymer surfaces onto which antibacterial small molecules or macromolecules were grafted, were either bulk materials or coatings themselves; in the latter approach, the polymeric coating serves as an adhesive interlayer and provides the chemical surface groups, for covalent grafting, that are not available on the underlying bulk material/device. Antimicrobial agents that contain chemically reactive groups such as hydroxyl, carboxyl, amino, etc., can be covalently linked to a wide variety of polymer surfaces using well-known facile chemical interfacial reactions, but those that do not possess convenient chemical groups for interfacial covalent bonding, such as furanones (see Section 11.4.4), can be linked using less common chemical strategies. In the following sections, we discuss selected representative approaches that illustrate examples of strategies used in terms of antibiotics and linking reactions.

11.4.1 Quaternary ammonium compounds Surfaces equipped with quaternary ammonium compounds (QACs) have attracted considerable interest [38]. QACs are antiseptics effective against both Gram-positive and Gram-negative bacteria. The mechanism of antibacterial action of QACs is still not fully resolved; however, two hypotheses appear to be lead contenders. The first and most-cited hypothesis is that sufficiently long cationic polymer chains penetrate the cell membrane [5, 38]. The second hypothesis, proposed by Kugler et al. [39], proposes that a highly charged surface can induce ion exchange between the positive charges on the surface and structurally essential mobile cations within the membrane. Being in close contact with a cationic surface, vital divalent cations cannot perform their normal role in charge neutralization of the head groups of membrane lipids, which results in a loss of membrane integrity. In an elegant set of experiments Murata et al. [40] investigated both hypotheses, varying the chain length containing the QA group and the surface density of QA groups in a gradient manner. It was concluded that the density of surface QA groups, and thus the density of cationic surface charges, is a key parameter, but it was not excluded that membrane insertion of alkyl chains may also be a possible mechanism of action. The relative contributions of these modes may well differ with the molecular nature of the QAC. Chen et al. [41] have recently reported data

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suggesting that nanometre-sized holes are caused in living cell membranes by cationic nanoparticles. An extensive body of literature exists on the development of various QACs and methods for their application; we discuss here only a limited number of representative studies. Silane based QAC have received considerable attention because of the feasibility of immobilisation onto hydroxylated surfaces such as metal oxides on orthopaedic implant surfaces [42±46]. The self-assembly of silanes on hydrophilic surfaces is well known; the silane group anchors covalently to the substrate and the quaternary ammonium group remains available for bioactivity. Covalent attachment of silane-QACs is feasible onto a variety of substrates including glass, cotton, and ceramics, as well as onto polymers such as silicon rubber after plasma activation to introduce hydroxyl groups [42]. Covalent bonding of QACs to substrates may lead to long-term antibacterial activity as it avoids leaching of the antibacterial compound over time but this has not been clinically verified yet. Silanes do have some drawbacks, and a number of alternative methods have been investigated for the generation of QAC surfaces. Plasma polymerisation cannot achieve a substantial QAC surface in one step but the deposition of an amine plasma polymer followed by quaternisation with hexyl bromide [9] or a 4vinyl-pyridine plasma polymer quaternised with bromobutane [47] leads to antibacterially active surfaces. In another study [48], plasma polymerisation of propionaldehyde was used to generate surfaces bearing aldehyde groups onto which amine-terminated starburst dendrimers were covalently immobilised by reductive amination. Remaining amine groups were then quaternised, which was verified by the emergence in XPS in signals assignable to positively charged amine groups at > 401.5 eV binding energy. Testing in vitro, however, showed a reduction in bacterial attachment of only 50%, which was deemed insufficient for practical purposes. Analogously, other amine macromolecules can be grafted onto plasma aldehyde or carboxylic acid surfaces and then quaternised. Graft polymerisation of vinylic or acrylic monomers onto plasma pre-treated surfaces followed by quaternisation is another option [49]. As for silver releasing coatings, test results show that some QACs on surfaces can be strongly antibacterial in vitro. There are, however, concerns related to the use of QAC on medical devices [45]. The main concern remains the cytotoxicity of QACs. It stands to reason that membrane disruption by QACs, whether by insertion or an ion disruption mechanism (or both), would equally apply to human cell membranes. For example, a recent study showed that an antimicrobial composite of poly(4-vinylpyridine)-co-poly(vinyl-N-hexylpyridinium bromide) and AgBr nanoparticles caused disruption and activation of blood platelets, rendering this material unsuitable for blood contacting applications [31], though it is not clear whether it is the cationic copolymer or the Ag‡, or both, that caused the adverse effects. Analogous to silver, however, in some

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clinical implant applications local disruption and cell death may be tolerable, whereas, for example, for contact lenses the irritation caused by QACs on the eye and the eye lid would likely not be tolerable.

11.4.2 Other cationic compounds A number of synthetic or natural macromolecules have been reported to be bactericidal [50±54]; they typically are cationic and therefore comparable to QACs. Cationic peptides such as melittin are well known for their antibacterial action but they often also damage human cell membranes. Several studies have made use of the natural biopolymer chitosan, which may be better tolerated by the human host body, but the reductions conferred by it to bacterial growth and biofilm formation appear to be less marked than for some other approaches. Recent examples of chitosan grafted onto surfaces following their activation by corona or air plasma are found in [55±57]. Lysozyme is a cationic protein present in the human tear fluid to help protect the eye against bacterial infections. Conte et al. [58] immobilised lysozyme onto plasma-treated polyethylene and showed antibacterial activity that varied with plasma treatment conditions and the concentration of lysozyme. Thus, a wide variety of cationic surfaces have been found to possess antibacterial activity in vitro, but their utility for human clinical usage remains to be investigated in more detail before one can confidently decide which applications/products may be suitable for such coatings. As for QACs, however, a key question is whether the cationic properties needed for effective action can be tolerated by human host tissue, or whether any damage to cells and tissue is relevant.

11.4.3 Established antibiotics The provision of known antibiotics such as penicillin and vancomycin on the surface of biomedical devices is an obvious route towards tackling the infection problem. However, it is not all that straightforward to design and implement effective coatings with predictable structures. Adsorptive (non-covalent) application of such low molecular weight compounds by themselves is not a desirable option, as such relatively low molecular weight compounds when physisorbed on a biomedical device surface would likely be displaced quite rapidly in biological media, with consequent expected rapid loss of antibacterial action. Thus, literature reports often described the conjugation of antibiotics onto polymer chains and then the ensemble molecules were coated adsorptively from solution onto materials surfaces. An alternative, much less widely used approach is the covalent grafting of antibiotics onto materials surfaces. The covalent grafting of antibiotic compounds onto medical device surfaces may provide a route for achieving permanent protection against biofilm

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formation without the detrimental effects associated with QACs. Several established, commercially available antibiotics have been grafted as (sub-)monolayer surface coatings, immobilised via covalent bonding, on various substrate materials. Various interfacial linking chemistries have been employed for that purpose, selected on the basis of reactive chemical groups available in the chemical structure of the antibiotic and on the substrate surface. The covalent surface immobilisation of antibiotics thus is related closely to the much larger body of work on covalent interfacial immobilisation of proteins as biomaterials coatings (e.g., [2] and Chapter 4). Analogous interfacial reaction schemes can be used, although antibiotics may not offer the convenient amine and carboxylic acid groups used prevalently for protein linking to surfaces via carbodiimide chemistry. Compared with the very extensive literature on covalent grafting of proteins, however, only a relatively small number of reports describe the covalent grafting of antibiotics, and only a few studies have used plasma-based approaches. Yet, they suffice to illustrate principles and challenges. Typically, plasma technologies are of greatest interest when the substrate material needs to be equipped with a suitable reactive chemical group. For chemically inert materials such as poly(tetrafluoroethylene), plasma surface treatments and plasma polymerisation are highly suited for producing a surface chemistry suitable for the covalent grafting of a (sub)monolayer of an antibiotic molecule. Aumsuwan et al. [59] used a microwave plasma of maleic anhydride, which upon hydrolysis formed carboxylic acid groups onto which penicillin could be grafted via a polyethylene glycol spacer. Such strategies can be designed to interface with different surface engineering strategies; for example Wach et al. used an analogous approach to construct vancomycin on PEG graft coatings on activated ceramic substrates [60]. Earlier work, also on titanium substrate, suggests that the PEG linker is not necessary for antibacterial activity of vancomycin [61]. A further example is the work by Ys et al. [62] in which Novobiocin was grafted onto amine surfaces by a Michael-type addition reaction of a distal C=C bond on Novobiocin, via oxymercuration reaction. Amine surfaces were prepared in two ways: either by the plasma polymerisation of n-heptylamine, or by plasma polymerisation of propanal (propionaldehyde) followed by solution grafting of polyallylamine. Both approaches were successful on a number of substrates, from silicon wafers (for spectroscopic analyses) to Teflon FEP sheet material. Sequential layers were characterised by XPS and ToF-SIMS. Incubation of samples with broth containing Staphylococcus epidermidis showed marked reduction in bacterial attachment (Plate V between pages 208 and 209). It is thought that the residual attachment of bacteria was at least in part due to coating defects that were unavoidable under the conditions used. Dust particles temporarily or permanently adhered to the surface at various steps, as the coating steps were not performed under clean room conditions, can cause

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such small defects, and indeed some dust particles were observed in optical microscopy. Nevertheless, a substantial reduction in bacterial attachment is clearly evident. In such work, a number of issues need to be considered. First, the covalent grafting may cause side reactions that may convert the molecules to a different molecular structure; for example, -lactams and furanones readily undergo ring opening in alkaline and acid media; thus, grafting must be performed in close to neutral pH conditions. Second, covalently grafted molecules may have geometrically restricted ability to perform their function, such as integrin docking or membrane lysis. Third, some antibiotics are thought to have an intracellular action, such as the gyrase inhibition activity of Novobiocin and the lux gene deactivation by furanones; can such compounds be active when covalently attached onto a medical implant and thus unable to diffuse into bacterial cells? Or do they act in different ways when in solution and grafted onto a material, respectively?

11.4.4 Furanones Many antibiotics are derived from leads provided by compounds extracted from natural sources, and plants and marine organisms continue to provide interesting lead compounds. Antimicrobial compounds extracted from natural sources form the basis of many novel therapeutic agents. Defence mechanisms against microbial colonisation, commonly through the biosynthesis of antimicrobial compounds, are observed in a variety of organisms, from marine life to terrestrial plants, and halogenated furanones (also known as fimbrolides), represent an interesting example [63]. They were extracted from an Australian marine algae, Delisea pulchra, and found to have strong antibacterial activity in solution against medically relevant bacterial strains [64]. The antibacterial nature of furanones was attributed to their ability to interfere with two bacterial cell communication strategies ± quorum sensing [65] and swarming [66]. The question then arose whether furanones would also be active when grafted onto a solid biomaterials surface. The use of furanones to prevent implant infection would offer several advantages; such as absence of toxicity of some furanones to fibroblasts [67]. Moreover, furanones are thought unlikely to lead to resistant strains of bacteria if the bacteria are not killed by furanones but biofilm formation inhibited by interference with quorum sensing, which is an essential step in the phenotype transformation of surface-adhered bacteria progressing to biofilm formation. Furanone compounds of interest are based on the 4-bromo-5-methylene2(5H)-furanone structure (Fig. 11.2). The furanone ring structure and the substituent in position R2 appear to be essential for antibacterial function, and the molecule thus needs to be covalently attached to surfaces via a reactive functional group remote from the furanone ring, such as R1 or the end of the

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11.2 Chemical structures of furanones investigated as surface-grafted antibacterials. From [68].

alkyl side chain. From a synthesis perspective, incorporation of a suitable group at R1 is the easier option although it is not straightforward. However, solution testing showed that the most active compounds were those where R1=H and activity also varied with the length of the alkyl chain. The absence of convenient chemical groups on available furanones led Muir et al. [69] and Al-Bataineh et al. [68, 70±72] to develop nitrene-based approaches for the covalent immobilisation of furanones onto surfaces of model biomedical device materials. Photo-chemical activation of surface azide groups was used to generate reactive nitrene radicals that formed covalent bonds with pre-adsorbed furanone molecules. Covalent attachment was confirmed by resistance to extensive washing. Nitrene radicals react indiscriminately with various bonds; their reaction with pre-adsorbed furanone molecules has little chemical specificity and will thus attach an unknown percentage of molecules in ways that interfere with activity (e.g., by linking into R2), though the alkyl tail should provide a sufficient number of linking sites based on statistical distribution of nitrene radical attack. Al-Bataineh et al. [68] developed two multilayer strategies for the grafting of furanones via nitrene surface groups. In one strategy, azide surface groups were attached onto polyacrylic acid graft layers and in the other onto PEG interlayers; in both cases a first layer of an n-heptylamine plasma polymer coating served as a platform for the covalent bonding of these reactive hydrogel graft layers (Fig. 11.3). The plasma polymerisation approach was chosen to enable ready transfer of these coating methodologies onto a variety of biomedical devices. Surface spectroscopic evidence showed that both strategies were effective in achieving a high density of surface-bound furanone molecules. The furanone multilayer coatings were extensively characterised by X-ray photoelectron spectroscopy [68, 70, 71] and Time-of-Flight Secondary Ion Mass Spectrometry [72] to probe for molecules attached to the surfaces and to check that functional groups important for antibacterial activity (the lactone ring and the bromines) were still present (though it was not feasible to detect molecules damaged by nitrene attack). Although it is possible that a percentage of the grafted molecules were attached such as to be inactive, biological assays showed considerable antibacterial activity of nitrene-grafted furanones against S. aureus with 75±90% reductions in growth and biofilm formation even when samples

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11.3 Multilayer grafting scheme for the covalent immobilisation of furanones. From [68]

were pre-coated with proteins [68]. A sufficient percentage of the molecules must have been attached in a way that did not interfere with their biological function. To avoid possible problems arising from radical reactions attacking the contact lens substrate, Zhu et al. [73] immobilised furanones onto silicone hydrogel contact lenses via a Michael-type addition reaction using an acrylate

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11.4 Schematic diagram of the attachment strategy for furanone-coated contact lenses using amine groups attached by grafting poly(allylamine) onto an aldehyde plasma functionalised contact lens surface; from [73].

group in position R1, which could be reacted with a surface amine group. This strategy is advantageous in terms of the linking reaction being specific to the acrylate side group and thus prevention of destruction of the activity of some of the immobilised furanone molecules, but on the other hand the immobilisation reaction is much slower and the antibacterial activity of the furanone molecule decreases upon functionalisation with the acrylate group at R1. Amine surface groups were produced in two ways: by plasma polymerisation of n-heptylamine onto contact lenses, and by plasma polymerisation of propionaldehyde followed by solution grafting of poly(allylamine). Again the furanone molecules fully resisted removal by extended washing, indicating that covalent surface attachment had indeed occurred, as opposed to physical adsorption or diffusion into the substrate material. The latter strategy, shown in Fig. 11.4, was advantageous in providing a more hydrophilic character to the multilayer coating structure, and was therefore selected for application to contact lenses used for clinical trials. Contact lenses coated with furanones were studied in both guinea pigs, where the lenses were worn for 1 month, and human volunteers during 22-hour wear. Reduction in bacterial adhesion of up to 92% was observed when comparing furanone-coated lenses with control lenses in guinea pigs, and human volunteers were used to assess wear comfort, which was identical to that of control lenses [73]. While these results are promising, cytotoxicity and mutagenic properties of these compounds have not been fully established.

11.4.5 Serrulatanes Serrulatanes comprise a class of naturally occurring diterpenes bearing hydrophilic substituents; they represent another set of antimicrobial compounds derived from natural sources. Their immobilisation onto surfaces serves here to illustrate strategies and challenges. These compounds are produced by some Australian plants in the genus Eremophila, which attracted attention because of their extensive use in traditional Australian Aboriginal medicine. For example,

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11.5 Chemical structure of some antibacterial constituents of leaf resins of Eremophila neglecta.

leaves of the resin-producing Eremophila duttonii plant were applied to treat minor dermal wounds and infected lesions, while decoctions were used as a sore throat gargle and for treatment of eye and ear pain [74]. The symptoms treated with Eremophila plant extracts suggest antibacterial activity, and this was indeed found to be the case with extracts in solution [75, 76] including against methicillin-resistant Staphylococcus aureus. Isolation and structural elucidation of antimicrobially active compounds from several Eremophila species has led to the identification of novel serrulatanes that possess activity against bacteria associated with biomedical device infections [77, 78]. Figure 11.5 shows the serrulatane skeleton and the structures of two representative active serrulatanes isolated from Eremophila plants. In all cases of activity, the serrulatane diterpene hydrocarbon skeleton is modified with hydrophilic substituents such as ±COOH, ±OH, ±OMe or ±OAc in one or both rings. As for furanones, approaches based on plasma polymer interlayers were chosen for grafting serrulatanes onto material surfaces, in order to facilitate transfer of the strategy onto various materials of biomedical interest. Two interfacial coupling methods were implemented in order to access suitable reactive structures on serrulatanes [62]. One strategy targets the allylic double bond in the `tail' of the molecules; it can be reacted with amines in a Michaeltype addition reaction. However, due to the dimethyl substitution, the allylic double bond of the serrulatanes shown in Fig. 11.5 and related compounds has relatively low reactivity, and it was found that catalysis by oxymercuration was required in order to achieve adequate reaction rates. This strategy enables immobilisation of all active serrulatanes identified to date, as all possess this allylic double bond. The other strategy is restricted to compounds that possess a carboxylic acid group where water-soluble carbodiimide chemistry can be utilised for coupling to surface amine groups. Thus, for both scenarios, surface amine groups were required and were produced either by plasma polymerisation of n-heptylamine or by plasma polymerisation of propionaldehyde followed by solution grafting of poly-

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allylamine onto the surface aldehyde groups. Either approach leads to successful grafting of serrulatanes but the latter yields significantly more hydrophilic surfaces, which may be advantageous for some applications, such as contact lenses. For example when attaching 8-hydroxyserrulat-14-en-19-oic acid via carbodiimide mediated amide formation onto a polyallylamine interlayer (on a propionaldehyde plasma polymer layer), in vitro studies with a biofilm-forming S. Epidermidis strain showed a reduction in bacterial colonisation to > 99% over 2 and 4 hr periods, compared to control samples Optical micrographs of bacterial adherence and subsequent biofilm formation of amine plasma polymer control surfaces and serrulatane-coated surfaces are shown in Plate VI(a,b) and (c,d) (between pages 208 and 209), respectively. On the serrulatane-coated surface, no biofilm growth was observed up to 48 hrs (at which time the experiments were stopped, in the view that if biofilm does not form within 48 hrs, it will not do so later). The fact that the amine surface became heavily colonised with this aggressive biofilm forming strain while serrulatane-grafted surfaces remained entirely free of biofilm, suggests that these compounds are potential candidates for biomedical applications. Surface analyses by XPS and ToF-SIMS of furanones and serrulatanes covalently grafted as antimicrobial coatings illustrate the challenges in surface analysis of multilayer coatings functionalised with low molecular weight molecules, especially if they are present in low density on the surface and do not possess a unique marker that can be easily probed. Detailed surface analysis is, however, essential in order to verify that the intended coating chemistries and structures have indeed been achieved; unfortunately, many publications do not show sufficient surface analysis data, and some inconsistencies in biological responses may well be assignable to surfaces being other than those intended, or only partly so. In particular, side reactions/products are very difficult to detect at times, whereas competent surface analysts are familiar with the common surface contaminants of siloxanes, hydrocarbon, and partially oxidised hydrocarbon compositions; these can be excluded by appropriate protocols and clean work. Bromine in furanones provide a convenient marker for the presence of these molecules [68, 70, 71] and, via XPS elemental compositions and assuming a uniform layer, the Br % can be used to estimate surface coverage. Br ions can also be detected in the negative ToF-SIMS spectra [72]. The higher sensitivity enables detection of Br and hence furanones at lower surface coverage than is possible by XPS but quantification by ToF-SIMS must be approached with caution even for closely related samples. In another report [73], some non-brominated furanones were grafted, and those surfaces were much more challenging to analyse by XPS. A shake-up satellite in the XPS C 1s signal indicated that aromatic structures were present, but the weakness of this component resulted in low signal-to-noise. The same applied when grafting serrulatanes; again a shake-up satellite assignable to

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aromatic structures was observed but again at low intensity [62]. An increase in the neutral hydrocarbon component at 285.0 eV also was consistent with expectation when grafting a diterpene. Better evidence, however, was supplied by ToF-SIMS, though quantitation remains an issue. It was found that in the TOF-SIMS positive ion spectra there were signals of significant intensity assignable to serrulatane fragments, such as the aromatic ring structure, which is conducive to yielding higher mass ions with good intensity. It was even feasible to detect an ion containing the entire serrulatane molecule but with the ±OH of the ±COOH group replaced by an ±NH (plus a proton for generating a positive ion) as would be expected for carbodiimide mediated amide linking [62]. More intriguing, however, was that detailed interpretation of the many ion peaks revealed the presence of a side reaction product, an N-acyl-urea derivative of the serrulatane, formed presumably by rearrangement of O-acyl isourea to the N-acyl form [62]. These examples indicate the great value of detailed surface analysis in the development of antibacterial (and other) coatings, so that observed biological responses can be interpreted with confidence and reliability.

11.5

Future trends

The continuing problems arising from infection of implants and biomedical devices mandate the development of better antibacterial surfaces. Surface modification and coating technologies are ideally suited to addressing this challenge as they can be performed such that bulk material properties are not altered. A wide variety of approaches have been reported, based on one, or occasionally two, of the following three concepts: · resistance to adhesion (non-fouling surfaces/coatings); · release of an antibacterial compound or ion; · covalent grafting of an antibacterial compound. Each of these concepts has specific advantages and disadvantages. It is therefore essential that surfaces resistant to bacterial biofilm colonisation be developed with a specific biomedical application in mind, targeted to the specific requirements. One area for future research is to combine improved understanding of the local biological environment with rational design of antibacterial materials approaches. Non-fouling coatings have proven to be able to resist bacterial colonisation and are of great promise for some applications, such as some catheters. For other biomedical devices and implants, however, it is necessary to investigate by clinical trials whether the ability to resist adhesion of biological material will also lead to poor tissue integration of an implant. Silver (or, to be more precise, silver ions) will continue to attract attention because of proven effectiveness, but a challenge is to gain a better understanding

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of its effects on mammalian tissues and cells, and to which extent toxicity arising from silver ions can be tolerated in specific biological environments. Analogously, coatings of quaternary amine compounds can be highly effective against bacteria, but again it would seem highly desirable to gain better appreciation of any adverse effects. The importance and consequences of adverse effects is likely to vary considerably between biomedical devices. To reduce potential issues with adverse host cell and tissue reactions, antibacterial compounds are desirable that possess high activity against bacteria but cytotoxic effects on human cells should occur only at much higher concentrations, if at all. The identification of promising antibiotics that can be applied as surface coatings is an important area of research towards infectionresistant biomaterials. It is also necessary that coatings are well characterised so that ensuing biological responses can be interpreted with confidence. The literature on antibacterial coatings and surfaces is growing rapidly, in response to increased awareness of the problem. The combination of biomaterials science, microbiology, and clinical science promises to lead to much improved outcomes.

11.6 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14. 15. 16. 17. 18.

References

P.C. Nicolson et al., US Patent US 6,951,894 (2005). K.S. Siow, L. Britcher, S. Kumar, H.J. Griesser, Plasma Proc. Polym. 3, 392 (2006). J.W. Costerton, P.S. Stewart, E.P. Greenberg, Science 284, 1318 (1999). R.O. Darrouiche, New England Journal of Medicine 350, 1422 (2004). L.G. Harris, R.G. Richards, Injury 37, S3 (2006). B.D. Hoyle, J.W. Costerton, Progress Drug Research 37, 91 (1991). R.C. Chatelier, X. Xie, T.R. Gengenbach, H.J. Griesser, Langmuir 11, 2576 (1995). R.C. Chatelier, X. Xie, T.R. Gengenbach, H.J. Griesser, Langmuir 11, 2585 (1995). S.N. Jampala, M. Sarmadi, E.B. Somers, A.C.L. Wong, F.S. Denes, Langmuir 24, 8583±8591 (2008). C. Susut, R.B. Timmons, Int. J. Pharm. 288, 253±261 (2005). K. Vasilev, P. Martinek, H.J. Griesser, J. Phys. Chem. Letters, submitted C.S. Kwok, T.A. Horbett, B.D. Ratner, J. Control. Release 62, 301±311 (1999). P. Kingshott H. Thissen, H.J. Griesser, Biomaterials 23, 2043±2056 (2002). P. Kingshott, S. McArthur, H. Thissen, D.G. Castner, H.J. Griesser, Biomaterials 23 4775±4785 (2002). P. Hamilton-Brown, T. Gengenbach, H.J. Griesser, L. Meagher, Langmuir 25, 9149± 9156 (2009). H. Thissen, T. Gengenbach, R. du Toit, D. Sweeney, P. Kingshott, H.J. Griesser, L. Meagher, Biomaterials 31, S510±S519 (2010). K.D. Park, Y.S. Kim, D.K. Han, Y.H. Kim, E.H. Lee, H. Suh, K.S. Choi, Biomaterials 19, 851±859 (1998). I. Fundeanu, H.C. van der Mei, A.J. Schouten, H.J. Busscher, Coll Surf B: Biointerfaces 64, 297±301 (2008).

ß Woodhead Publishing Limited, 2011

306

Surface modification of biomaterials

19. L.G. Harris, S. Tosatti, M. Wieland, M. Textor, R.G. Richards, Biomaterials 25, 4135±4148 (2004). 20. P. Kingshott, J. Wei, D. Bagge-Ravn, N. Gadegaard, L. Gram, Langmuir 19, 6912± 6921 (2003). 21. M. Rai, A. Yadav, A. Gade, Biotechnol. Adv. 27, 76 (2009). 22. P. Favia, M. Vulpio, R. Marino, R. D'Agostino, R.P. Mota, M. Catalano, Plasmas Polym. 5, 1±14 (2000). 23. E. Sardella, P. Favia, R. Gristina, M. Nardulli, R. D'Agostino, Plasma Process Polym. 3, 456±469 (2006). 24. B. Despax, P. Raynaud, Plasma Process Polym. 4, 127 (2007). 25. K. Vasilev, V. Sah, K. Anselme, C. Ndi, M. Mateescu, B. Dollmann, P. Martinek, H. Ys, L. Ploux, H.J. Griesser, NanoLetters 10, 202±207 (2010). 26. N. Poulter, X. Munoz-Berbel, A.L. Johnson, A.J. Dowling, N. Waterfield, A.T.A. Jenkins, Chem. Comm. 2009, 7312±7314. 27. V. Zaporojtchenko, R. Podschun, U. SchuÈrmann, A. Kulkarni, F. Faupel, Nanotechnology 17, 4904±4908 (2006). 28. J.E. Gray, P.R. Norton, R. Alnouno, C.L. Marolda, M.A. Valvano, K. Griffiths, Biomaterials 24, 2759±2765 (2003). 29. K. Vasilev, J. Cook, H.J. Griesser, Expert Review of Medical Devices 6, 553±567 (2009). 30. A.M. Schrand, L.K. Braydich-Stolle, J.J. Schlager, L. Dai, S.M. Hussain, Nanotechnology 19, 235104 (2008). 31. K.N. Stevens, M.L. Knetsch, A. Sen, V. Sambhy, L.H. Koole, ACS Appl. Mater. Interf. 1, 2049±2054 (2009). 32. S.S. Griesser, K. Vasilev, H. Ys, C. Ndi and H.J. Griesser, 20th Annual Meeting of the Australasian Soc Biomater and Tissue Eng., Brisbane, Australia, 10±12 Feb. (2010). 33. D.K. Riley, D.C. Classen, L.E. Stevens, J.P. Burke, Amer J Med 98, 349±356 (1995). 34. A. Srinivasan, T. Karchmer, A. Richards, X. Song, T.M. Perl, Infection Control and Hospital Epidemiology 27, 38±43 (2006). 35. A. Daniel, C. le Pen, C. Archambeau, F. Reniers, Appl. Surf. Sci. 256, S82±S85 (2009). 36. W. Zhang, Y.H. Zhang, J.H. Ji, Q. Yan, A.P. Huang, P.K. Chul, J. Biomed. Mater. Res. Pt A 83A, 838±844 (2007). 37. Y.M. Jeong, J.K. Lee, S.C. Ha, S.H. Kim, Thin Solid Films 517, 2855±2858 (2009). 38. J.C. Tiller, C.J. Liao, K. Lewis, A.M. Klibanov, Proc. Natl. Acad. Sci. U.S.A., 98, 5981 (2001). 39. R. Kugler, O. Bouloussa, F. Rondelez, Microbiology-Sgm 151, 1341 (2005). 40. H. Murata, R.R. Koepsel, K. Matyjaszewski, A.J. Russell, Biomaterials 28, 4870 (2007). 41. J.M. Chen, J.A. Hessler, K. Putchakayala, B.K. Panama, D.P. Khan, S. Hong, D.G. Mullen, S.C. DiMaggio, A. Som, G.N. Tew, A.N. Lopatian, J.R. Baker, M.M.B. Holl, B.G. Orr, J. Phys. Chem. B 113, 11179±11185 (2009). 42. B. Gottenbos, H.C. van der Mei, F. Klatter, P. Nieuwenhuis, H.J. Busscher, Biomaterials 23, 1417 (2002). 43. J.Y. Huang, R.R. Koepsel, H. Murata, W. Wu, S.B. Lee, T. Kowalewski, A.J. Russell, K. Matyjaszewski, Langmuir 24, 6785 (2008). 44. O. Bouloussa, F. Rondelez, V. Semetey, Chem. Comm. 2008, 951. 45. R.A. Brizzolara, D.M. Stamper, Surf. Interf. Analysis 39, 559 (2007). 46. M.J. Saif, J. Anwar, M.A. Munawar, Langmuir 25, 377 (2009).

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47. W.C.E. Schofield, J.P.S. Badyal, ACS Appl. Mater. Interf. 1, 2763±2767 (2009). 48. D. Weber, N.R. Choudhury, H.J. Griesser, in preparation. 49. C. Yao, X.S. Li, K.G. Neoh, Z.L. Shi, E.T. Kang, Appl. Surf. Sci. 255, 3854±3858 (2009). 50. G. Cheng, H. Xue, Z. Zhang, S. Chen, S. Jiang, Angew. Chem. 120, 8963 (2008). 51. M.W. Huh, I.K. Kang, D.H. Lee, W.S. Kim, L.S. Park, K.E. Min, K.H. Seo, J. Appl. Polym. Sci. 81, 2769±2778 (2001). 52. B. Gottenbos, D.W. Grijpma, H.C. Van der Mei, J. Feijen, H.J. Busscher, J Mater. Sci.: Mater. Med. 10, 853 (1999). 53. T.P. Martin, S.E. Kooi, S.H. Chang, K.L. Sedransk, K.K. Gleason, Biomaterials 28, 909 (2007). 54. E.I. Rabea, M.E.T. Badawy, C.V. Stevens, G. Smagghe, W. Steurbaut, Biomacromolecules 4, 1457 (2003). 55. R.D. Joerger, S. Sabesan, D. Visioli, D. Urian, M.C. Joerger, Packaging Tech. Sci. 22, 125±138 (2009). 56. H.J. Tseng, S.H. Hsu, M.W. Wu, T.H. Hsueh, P.C. Tu, Fibers Polym. 10, 53±59 (2009). 57. E.S. Abdou, S.S. Elkholy, M.Z. Elsabee, E. Mohamed, J. Appl. Polym. Sci. 108, 2290±2296 (2008). 58. A. Conte, G.G. Buonocore, M. Sinigaglia, L.C. Lopez, P. Favia, R. D'Agostino, M.A. Del Nobile, J. Food Protection 71, 119±125 (2008). 59. N. Aumsuwan, S. Heinhorst, M.W. Urban, Biomacromolecules 8, 3525 (2007). 60. J.-Y. Wach, S. Bonazzi, K. Gademann, Angew. Chem. Int. Ed. 47, 7123±7126 (2008). 61. J. Parvizi, E. Wickstrom, A.R. Zeiger, C.S. Adams, I.M. Shapiro, J.J. Purtill, P.F. Sharkey, W.J. Hozack, R.H. Rothman, N.J. Hickok, Clin. Orthopaedics Rel. Res. 429, 33±38 (2004). 62. H. Ys, Antibacterial coatings for biomedical devices by covalent grafting of serrulatane diterpenes, PhD Thesis, University of South Australia (2010). 63. R. de Nys, M. Givskov, N. Kumar, Antifouling Compounds, 42, 55±86 (2006). 64. M. Manefield, R. de Nys, N. Kumar, R. Read, M. Givskov, P. Steinberg, S.A. Kjelleberg, Microbiology ± UK 145, 283±291 (1999). 65. D. Ren, J.J. Sims and T.K. Wood, Environ. Microbiol. 3, 731±736 (2001). 66. T.B. Rasmussen, M. Manefield, J.B. Andersen, L. Eberl, U. Anthoni, C. Christophersen, P. Steinberg, S. Kjelleberg, M. Givskov, Microbiology ± UK 146, 3237±3244 (2000). 67. J.K. Baveja, M.D.P. Willcox, E.B.H. Hume, N. Kumar, R. Odell, L.A. PooleWarren, Biomaterials 25, 5003±5012 (2004). 68. S.A. Al-Bataineh, Spectroscopic characterisation of surface-immobilised antibacterial furanone coatings, PhD Thesis, University of South Australia (2006). 69. B.W. Muir, H. Thissen, R.W. Read, M.D.P. Wilcox, P. Erickson, H.J. Griesser, Proceeding of the 6th World Biomaterial Congress, Hawaii, USA (2000). 70. S.A. Al-Bataineh, L.G. Britcher, H.J. Griesser, Surf. Sci. 600, 952 (2006). 71. S.A. Al-Bataineh, L.G. Britcher, H.J. Griesser, Surf. Interf. Anal., 38, 1512 (2006). 72. S.A. Al-Bataineh, M. Jasieniak, L. Britcher, H.J. Griesser, Anal. Chem. 80, 430±436 (2008). 73. H. Zhu, A. Kumar, J. Ozkan, R. Bandara, A, Ding, I. Perera, P. Steinberg, N. Kumar, W. Lao, S.S. Griesser, L. Britcher, H.J. Griesser, M.D.P. Willcox, Optometry & Vision Sci. 85, 292±300 (2008). 74. E.L. Ghisalberti, J. Ethnopharmacol., 44, 1±9 (1994).

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75. E.A. Palombo, S.J. Semple, J. Ethnopharmacol., 77, 151±157 (2001). 76. C.P. Ndi, S.J. Semple, H.J. Griesser, M.D. Barton, J. Basic Microbiol. 47, 158±164 (2007). 77. C.P. Ndi, S.J. Semple, H.J. Griesser, S.M. Pyke, M.D. Barton, J. Natural Prod., 70, 1439±1443 (2007). 78. C.P. Ndi, S.J. Semple, H.J. Griesser, S.M. Pyke, M.D. Barton, Phytochemistry, 68, 2684±2690 (2007).

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Plate V Bacterial attachment after 4 hrs on surfaces of polyallylamine `control' (a) and Novobiocin grafted onto polyallylamine (b). From Ys H (2010) `Antibacterial coatings for biomedical devices by covalent grafting of serrulatane diterpenes', PhD Thesis, University of South Australia.

Plate VI Optical micrographs of bacterial adherence and subsequent biofilm formation on (a,b) amine plasma polymer control surfaces, and (c,d) serrulatane-coated surfaces. The diffuse colour in d) is due to diffusion of dye into the polymer material, not to biofilm staining as in b). From Ys H (2010) `Antibacterial coatings for biomedical devices by covalent grafting of serrulatane diterpenes', PhD Thesis, University of South Australia.

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Modifying biomaterial surfaces to optimise interactions with soft tissues J . G O U G H , University of Manchester, UK

Abstract: This chapter focuses on surface modification for soft tissue applications where the type of tissue is discussed along with various methods of surface modification. Typically the type of modification is governed by the type of tissue architecture, for example surface modification of highly organised tissues such as ligament tend to focus on contact guidance techniques, whereas tissues where there is a high degree of cell-matrix or cell-cell interplay focus on extracellular matrix or other biomolecule modifications. Key words: soft tissue engineering, liver, skin, kidney, skeletal muscle, ligament, tendon.

12.1

Introduction

This chapter will focus on surface modification of biomaterials for a range of soft tissue applications, specifically liver, kidney, tendon/ligament, skeletal muscle and skin. The type of surface modification exploited will range from topography, protein adsorption (specific and non-specific), sugars, other biomolecules and chemistry. Surface modification is generally accepted as a route to enhanced cell functions for biomaterials and tissue engineering applications. Consideration must be given to modifying cell morphology to enhance cell behaviour, density or orientation of the modifier and stability of the modifier on the surface. Clearly the type of modification will be dictated by the specific cell niche and architecture of the tissue in question. In addition, the specific chemistry of the variety of materials used will dictate the ability to covalently couple various groups and biomolecules to gain the enhanced cell behaviours required.

12.2

Surface modification of biomaterials for the liver

The liver is one of the largest and most complex (in terms of structure and function) organs of the body and contains several different cell types as well as an extensive blood supply. The cuboidal hepatocyte is the main functional cell

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of the liver and therefore the majority of studies involving biomaterials and tissue engineering have focussed on this cell. However, it is becoming increasingly apparent that consideration must be given to other vital cells of the liver such as the hepatic stellate cell. Biomaterials are being developed for potential tissue engineering of the liver but also for extracorporeal bioartificial liver devices to support patients with liver failure whilst waiting for liver transplantation. These systems are intended to be similar to kidney dialysis for the processing of blood/plasma. In vitro alternatives to animal testing are also being developed as liver mimics to enable chemical/pharmaceutical toxicity testing, drug development and the study of drug metabolism, for example. Developing tissue engineering strategies for liver is a major challenge due to the difficulty in the growth and maintenance of hepatocytes in vitro. A number of strategies have been developed, some primarily to improve the in vitro culture of these cells, but which therefore have applications to tissue engineering. While hepatocytes can be cultured in vitro on a number of substrates, their spreading and therefore deviation from the in vivo morphology has been found to correlate with dedifferentiation and loss of hepatocyte-specific function (see LeCluyse et al., 1996 for a comprehensive early review), even when a variety of extracellular matrix molecules have been used as the substrate.

12.2.1 Extracellular matrix molecules Liver cells interact with several ECM proteins including collagens I-IV, fibronectin, laminin and heparin sulphate proteoglycans, therefore there has been much research dedicated to simple adsorption of some of these ECM components with the assumption that this will increase the hepatocyte viability and function in vitro. This has not been the case, however, due to rapid dedifferentiation. For example, Hodgkinson et al. (2000) demonstrated that adult male rat hepatocytes lost CYP2C11 expression after fibronectin induced cell spreading. Sudhakaran et al. (1986) also noted dedifferentiation markers (decreased albumin synthesis and increased -fetoprotein synthesis) were more evident on fibronectin compared to collagen IV. Decreases in albumin and many P450 enzymes alongside increased expression of cytoskeletal markers, organisation and spreading seem to occur on many ECM-modified 2D substrates. However adaptation of culture methods may improve this. For example, culture of rat hepatocytes on poly(lactic-coglycolic) acid fibrous matrices coated with either collagen I, IV, laminin or fibronectin (adsorbed from 3 g/ml solutions) resulted in an increase in number and secretion of more albumin compared to untreated control matrices (Fiegel et al., 2004).

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12.2.2 Plasma treatment Microchannel scaffolds can be utilised to direct cell migration, proliferation and behaviour and combined with surface modification for directed attachment and differentiation, may be useful in promoting the generation of specific tissues in 3D. Plasma-treated rectangular polydimethylsiloxane (PDMS) microchannels have been fabricated for the growth of liver explants from chick embryos (and also kidney as detailed later) (Leclerc et al., 2008). Air plasma treatment (15, 40 and 900 seconds) was used to modify the usually hydrophobic (106) nature of PDMS. Treated PDMS contact angles changed to approximately 40  6; 56  4; 67  4 and 69  4 after 24, 48, 72 and 96 h respectively. This resulted in faster cell migration through the microchannels, therefore suggesting a benefit of both microchannels and surface modification to enhance cell migration through 3D porous tissue engineering scaffolds.

12.2.3 Sugars Lactose and heparin have been used to modify a polystyrene-based foam for rat hepatocyte culture. Amino derivatised foams were prepared and lactose (oxidised to potassium lactonate) and heparin dissolved in 0.1 M MES were added along with the coupling reagent 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride. The modifications resulted in enhanced attachment and testosterone metabolism via cytochrome P450 compared to adsorbed collagen substrates. Albumin secretion was also demonstrated up to 7 days (Gutsche et al., 1996). Galactose has been heavily exploited as a surface modification for hepatocyte culture and liver tissue engineering (reviewed in Cho et al., 2006) due to the presence of the galactose (and N-acetylgalactosamine)-binding asialoglycoprotein receptor (ASGPR) on the hepatocyte cell surface. Galactosylated chitosan/alginate sponges were formed by freeze drying after forming galactosylated chitosan from chitosan and lactobionic acid. Increased hepatocyte functions in terms of spheroid formation, cell-cell contacts, albumin secretion and p450 activity were observed, especially when co-cultured with fibroblasts (Fig. 12.1) (Seo et al., 2006). Weigel et al. (1978) discovered specific binding of rat hepatocytes to galactosides immobilised onto polyacrylamide gels, and later found that this binding was via the asialoglycoprotein receptors (Weigel, 1980). Lactose-substituted polystyrene (poly-N-p-vinylbenzyl-D-lactonamide), was also found to bind primary rat hepatocytes via their ASGPRs (Tobe et al., 1992) and in this case the cells were able to retain a spherical shape and form multilayer aggregates with the addition of EGF and insulin. This spheroid formation resulted in longer viability and albumin secretion, more indicative of hepatocytes in vivo.

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12.1 Micrographs of hepatocytes stained with MTT reagent within alginate (a) and alginate/galactose (b) after 36 hours of culture. RT-PCR analyses (c) of Cx32 and E-cadherin after 12, 36 and 60 hours of culture. Reprinted from Seo et al., `Enhanced liver functions of hepatocytes cocultured with NIH 3T3 in the alginate/galactosylated chitosan scaffold', Biomaterials, 27, p. 1491. Copyright (2006), with permission from Elsevier.

Other factors have been found to influence hepatocyte function such as coculture, composite systems, 3D encapsulation but are beyond the scope of this book.

12.3

Surface modification of biomaterials for the kidney

Biomaterials for kidney applications include tissue engineering strategies but also and possibly more commonly, development of new bioartificial kidneys for improved dialysis. The need for new kidney treatment strategies is on the increase due to the increase in people suffering from diabetes and obesity. Cellular requirements for kidney cells are likely to be similar to those of hepatocytes based on the complex architecture of the kidney. For bioartificial kidneys, modification tends to be in the form of matrix biomolecules on appropriate polymer films. For tissue engineering, the 3D scaffold architecture along with bioactive modification is important.

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12.2 Micrographs of LLCPK-1 cells on poly(L-lactide-co-caprolactone) microspheres. (a): 20, (b): 40, (c): 20 P-15 modified, (d): 40 P-15 modified. 3D arrangements of cells can be seen around the microspheres. Reprinted from Garkhal et al., `Surface modified poly( L -lactide-cocaprolactone) microspheres as scaffolds for tissue engineering'. Journal of Biomedical Materials Research, 2007, with permission from Wiley.

The plasma-treated rectangular PDMS microchannels described above for liver were also investigated for kidney cell adhesion and migration. Plasma treatment resulted in enhanced migration of kidney cells (from chick embryo kidney explants) and the migration was also triggered by fibronectin coating (Leclerc et al., 2008). Poly(L-lactide-co--caprolactone) microspheres have been modified with P15 (a 15 peptide synthetic analogue of collagen I bound to the microspheres using carbodiimide) (Garkhal et al., 2007, Fig. 12.2). Microspheres were initially treated with NaOH to generate free carboxyl end groups. Surface modified microspheres caused increased cell adhesion, proliferation and matrix formation in the cell line porcine proximal kidney tubule cells (LLC-PK1). Microspheres were observed to be interconnected via matrix formation. Pronectin F, collagen I or collagen IV have been used to modify porous filter membranes made of polysulfone or cellulose acetate. Attachment and proliferation of a porcine renal epithelial cell line (LLC-PK1) was higher on the polysulfone membranes than the cellulose acetate. Biomolecule coated membranes resulted in even higher attachment and proliferation. RT-PCR determined

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the expression of the sodium-glucose cotransporter 1 and glucose transporter 1 (Sato et al., 2005). Fissell et al. (2006) developed a microelectromechanical systems technology (MEMS) for the study of human renal tubule cells with a view to miniaturising current bioartificial kidney devices. Typical MEMS substrates (single-crystal silicon, polycrystalline silicon, silicon dioxide, silicon nitride, SU-8 photoresist) as well as a novel silicon nanopore membrane were studied, coated with collagen IV. Confluent monolayers with tight junctions and central cilia formed on the surfaces as shown by ZO-1 and acetylated -tubulin immunocytochemistry. As with liver tissue engineering, a critical step for kidney is to devise 3D systems capable of maintaining high levels of cell-cell contact. Due to the complicated architecture of the kidney, this particular organ is another major tissue engineering challenge.

12.4

Surface modification of biomaterials for tendons/ligaments

Tendon/ligament reconstruction is often required as a result of sporting injuries, commonly to the posterior or anterior cruciate ligaments, and less commonly the rotator cuff tendon. Biological grafts such as autografts from the patellar tendon or hamstring are the most commonly used repair strategies. Donor site morbidity and muscle atrophy are amongst the reasons for development of new tissue engineering approaches to repair. As with other tissues such as skeletal muscle, a great deal of attention has been on developing materials that mimic the architecture of the native tissue to provide a scaffold with the right biomechanical properties and also provide cues for guided tissue regeneration. Also studies have aimed at scaffolds that allow regeneration of a tissue structure that mimics that found in vivo (as shown in Fig. 12.3). Subsequent to these approaches has been surface modification to enhance cell attachment and behaviour. The ACL, for example, is a strong, highly organised collagen rich structure with additional elastin and proteoglycans. Fascicular organisation runs parallel along the long axis of the ligament. Natural healing and regeneration of the tissue is hampered by its lack of vasculature. Collagen 1 was used to micropattern surfaces prepared using PDMS microcontact printing (Fig. 12.4). Grooves were prepared of 10 m and 4 m height. Rat tail tenocytes were found to selectively attach and elongate along the long axis of the grooves when coated with collagen. Collagen at 1 mg/ml was found to significantly increase the proliferation rate of the tenocytes (Chen et al., 2008, Fig. 12.4). The naturally occurring protein silk has been extensively researched for ligament and tendon engineering due to its mechanical properties, hierarchical organisation similar to that of ligament tissue and long history of being a

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12.3 Haemotoxylin and eosin staining of (a) bMSC-knitted PLGA tissue engineered construct of Achilles tendon in a rabbit model at 24 weeks, and (b) normal Achilles tendon. Addition of fibrin glue to the construct resulted in faster tissue formation. Reprinted from Goh et al., 2003, with permission from Mary Ann Liebert Inc.

biocompatible material having been used in sutures for centuries (for review see, for example, Wang et al., 2006). Natural silks consist of fibroin, a filament core protein surrounded by a serecin coating. The most widely studied silks are those from the silkworm Bombyx mori and the spider Nephila clavipes. Serecin has been found to cause immune responses in vivo and therefore is generally removed for tissue engineering applications covering a wide range of tissues and organs. Although silk can be formed into many structures such as films, micro- and nanofibrous mats, hydrogels and porous sponges it has also been investigated as a coating for other materials (Chiarini et al., 2003). In this study, Chiarini et al. coated poly(carbonate) urethane (PCU) films with Bombyx mori silk fibroin and observed the effects on normal human fibroblasts. Twice as many cells were found to initially adhere compared to uncoated PCU. Total number of cells after 30 days of culture was also over two-fold. Higher metabolic rate (determined by glucose consumption) and higher collagen rich extracellular matrix was also found on the silk coated PCU. The RGD peptide sequence is found in silk from A.

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12.4 Scanning electron micrographs of tenocytes cultured on 10 m wide  4 m deep grooved PDMS substrates: (a) after 1 hour culture with collagen I modification, (b) after 24 hours with collagen I modification, (c) after 24 hours without modification and (d) after 24 hours on smooth PDMS control. Reprinted from Applied Surface Science, 2, Chen et al., Effects of micropatterned surfaces coated with type I collagen on the proliferation and morphology of tenocytes, p. 370. Copyright (2008), with permission from Elsevier.

pernyi but not from B. mori which has been found to result in increased cell adhesiveness to the A. pernyi silk (Minoura et al., 1995). Covalent coupling of RGD has subsequently been used to increase cell attachment to a variety of silks (Chen et al., 2003). Silk fibroin itself has also been used to modify the surface of a variety of polymers to enhance fibroblast attachment, proliferation and matrix production (for review, see Wang et al., 2006). Fibronectin has also been used to enhance cell responses. A study by Lu et al. (2005) determined that out of three resorbable polymer fibre-based scaffolds, PLLA-Fn scaffolds were better for primary rabbit ACL cells. The modification with Fn resulted in increased cell attachment and long term matrix production.

12.5

Surface modification of biomaterials for skeletal muscle

Regeneration of skeletal muscle is of clinical interest due to the need to rectify the loss of tissue due to cancer, trauma and congenital defects. For the development of biomaterials for tissue engineering of skeletal muscle, a topographical

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approach has often been utilised due the parallel arrangement of individual muscle fibres in vivo. During skeletal myogenesis, mononucleated fibroblasts fuse to form multinucleated myotubes that have a parallel arrangement.

12.5.1 Topographical Topgraphical cues for muscle cells have been investigated since the 1980s where Isaeva (1980) reported the orientation of myoblasts on substrates with grooves of 5±20 m. The effects of grooved substrata on fetal and neonatal myoblasts from BALB/ c mice have been investigated (Evans et al., 1999). These substrates were fabricated using photolithography on fused silica. Deeper grooves (2.3±6.0 m) resulted in alignment of both cell types. Fetal myoblasts also aligned in shallower grooves (0.04±0.14 m). The authors correlated this to the two-phase myogenesis process where primary myoblasts align to form a scaffold for secondary myotubes to align. The more developmentally mature myoblasts therefore respond to the larger grooves. Myoblasts on finer grooved surfaces (130 nm wide, 210 nm deep) coated with laminin showed a high level of alignment and migration was restricted to along the long axis of the grooves. Myotube formation also occurred along the long axis of the grooves (Clark et al., 2002). Studies in our laboratory on myoblasts grown on electrospun polymer scaffolds show preferential alignment and differentiation on oriented versus random fibre structures (Aviss et al., 2010). Lam et al. (2006, 2009) have reported alignment of myoblasts on laminin coated microtopgraphical featured PDMS substrates (Fig. 12.5). These myoblasts then differentiated into aligned myotube monolayers with greatest alignment on 6 and 12 m wave sizes.

12.5.2 Extracellular matrix molecules and their orientation While for some cell types there is a primary need for specific biomolecule interactions, for myoblasts the focus of extracellular matrix molecules has been on their patterning and orientation. Clearly the architecture of the extracellular matrix with particular reference to the fibrous components affects myoblast morphology. Oriented fibronectin deposition has also been shown to act as both a biological and topographical cue for myoblast alignment and aligned myotube formation (Turner et al., 1983). Oriented fibronectin adsorption was achieved by adsorbing fibronectin onto tissue culture plastic in a solution containing urea crystals which were then subsequently washed away. This results in areas of oriented fibronectin varying between 0.7 and 5 m in width. Stern et al. (2009) showed that muscle progenitor cells and C2C12 myoblasts cultured on muscle ECM extracts resulted in increased proliferation and

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12.5 Skeletal muscle tissue engineering: (a) diagram of construct formation of laminin coated wavy PDMS substrate, (b) optical micrograph of micro-sized waves (scale bar 50 m), (c) construct roll-up showing sutures and construct between (scale bar 6 mm), (d) primary rat muscle cells on flat (left) and waved (right) PDMS (scale bar 100 m), (e) cells on waved features inside (left) and outside (right) the gel (scale bar 100 m). Reprinted from Biomaterials, 30, Lam et al., Microfeature guided skeletal muscle tissue engineering for highly organized 3-dimensional free-standing constructs, p. 1152. Copyright (2009), with permission from Elsevier.

differentiation as measured by MyoD, myogenin and myosin heavy chain markers. Substrates coated with extracellular matrix components extracted from skeletal muscle increased proliferation of C2C12 myoblasts and human muscle progenitor cells compared to a coating of collagen alone. Myogenic differentiation was also determined using typical markers MyoD, myogenin and myosin heavy chain. Expression of myogenin and myosin heavy chain was enhanced on ECM-coated substrates. Analysis of myotube formation showed that the ECMcoated substrates showed larger myotubes with more nuclei/myotube suggesting the generation of more mature myotubes. Cell adhesive glass micropatterns of 5±100 m within non-adhesive organosilane treated glass allowed myoblast adhesion and alignment (Clark et al.,

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1997). Myoblasts were less well aligned on laminin (a component of the basal lamina and can promote differentiation) which bound to the organosilane treated micropatterns but the myoblasts preferentially attached to the laminin and alignment increased upon differentiation into myotubes. Laminin also promoted the long-term survival of these cells.

12.6

Surface modification of biomaterials for skin

Skin is the largest organ of the body and provides an essential barrier function to damage and infection. The need for skin regeneration and repair is therefore of utmost importance for patients suffering with burns, removal of tumours and chronic wounds. Materials are required that provide rapid wound closure, and promotion of tissue repair with a functional tissue and minimal scar formation.

12.6.1 Extracellular matrix proteins/peptides In some relatively early work, keratinocyte function has been assessed on tissue culture polystyrene coated with collagens I and IV, matrigel, RGD, fibronectin and vitronectin (Dawson et al., 1996). Matrigel and the two collagens were found to enhance keratinocyte attachment and proliferation. Individual amino acids coupled to cellulose via esterification have been investigated in their ability to direct cell behaviour. Fibroblasts were found to preferentially spread and adopt the usual spindle-like morphology on cellulose substrates modified with aromatic amino acids, in particular tryptophan (Kalaskar et al., 2008). It was also found that this increase in adhesion and spreading could be a result of increased fibronectin deposition on the surfaces, enhancing the cellular response. Integrin blocking experiments confirmed receptor mediated adhesion to fibronectin. Chitosan (a deacetylated derivative of chitin) is a polysaccharide which has been investigated for the potential regeneration of many tissue types. Chitosan has been modified with peptide sequences from laminin (Mochizuki et al., 2003, Masuda et al., 2009). Two of the sequences bind to cell surface heparin-like receptors (RLVSYNGIIFFLK and ASKAIQVFLLAG) promoted strong human keratinocyte attachment and were further investigated as keratinocyte delivery systems for a murine wound model (Masuda et al., 2009). These were found to promote keratinocyte migration and reduction in granulation tissue formation. Another peptide (EPDIM) has been immobilised to 500±1000 m diameter chitosan beads and been found to promote NIH 3T3 fibroblast and HaCaT keratinocyte proliferation (Bae et al., 2009) via integrin 3 1 interactions. Self-assembling monolayers (SAMs) based on alkanethiols of varying alkyl chain length and terminating group on gold have been used to study cell responses to defined chemistries. These model surfaces are very useful for the study of cell interaction with specific chemistries as very ordered monolayers

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are formed relatively easily and with high control over terminating group. 3T3 fibroblasts have been found to attach poorly to OH terminated SAMs whereas extensive attachment and spreading was observed on COOH terminated SAMs. Long chain CH3 terminated SAMs resulted in poor attachment and short chain CH3 terminated SAMs showed some attachment and spreading and was deemed intermediate (Cooper et al., 2000). The effect of fibronectin (Fn) conformation and concentration on human neonatal foreskin keratinocytes has also been investigated using similar SAM surfaces. The availability of the Fn synergy site, controlled by the density on the surface, affected integrin mediated attachment, spreading and differentiation. It was found that NH2 and CH3 (not the OH or COOH) terminated SAMs at saturated Fn density caused greater attachment and spreading. At low Fn surface density, OH terminated SAMs resulted in greater attachment and spreading. Differentiation assessed by involucrin staining suggested lower differentiation on these SAMs (Bush et al., 2008). Regarding fibroblasts, a relationship also exists between attachment and spreading in response to SAM terminating group.

12.6.2 Plasma polymerisation Plasma co-polymers of acrylic acid/octa-1,7-diene and allyl amine/octa-1,7diene have been shown to influence human keratinocytes. It was shown that surfaces containing low concentrations of COOH groups favoured keratinocyte attachment which was comparable to attachment to collagen I, a postive control for keratinocyte attachment. Plasma co-polymers containing nitrogen functional groups (mainly amine and imine) were also found to promote favourable attachment at higher concentrations but at a lower level compared to COOH functionalisation (France et al., 1998). The co-culture of melanocytes and keratinocytes has proven successful for repigmentation aswell as rapid re-epithelialisation for vitiligo patients. As with many co-culture systems it is a challenge to define conditions that are optimal for the desired response from both cell types. Beck et al. (2003) determined that the ideal balance between keratinocyte and melanocyte density was achieved by varying the culture media and amine functionalisation. Further studies have shown the use of a silicone carrier with 20% COOH functionalisation by plasma polymerisation as a useful system for the transfer of melanocytes and keratinocytes. PVC carriers were also studied but while attachment was good, transfer of cells was not. Silicone functionalised with 5% COOH or 100% allylamine was also inferior (Eves et al., 2008). The plasma polymerisation technology resulted in the development of `MySkinTM' ± PVC coated with plasma polymerised functional surface containing 20% carboxylic acid. It supports autologous keratinocytes and has been clinically effective in the treatment of burns and chronic foot ulcers. A

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weekly application of keratinocytes using this system allowed the gradual formation of granulation tissue followed by re-epithelialisation and full close of the ulcers as shown in Plate VII (between pages 208 and 209) (Haddow et al., 2003; Moustafa et al., 2004).

12.7

Conclusions

It seems that many studies involving surface modification of biomaterials for soft tissue applications focus on synthetic functional groups via, for example, gas plasma or simple biomolecule adsorption. Modifications of materials for tissues with specific ordered architecture (such as ligament) tend to focus on topographical cues in addition or alone. Plasma polymerisation has many advantages including the ability to deposit on most surfaces including those of variable topography and porosity. The distribution of the functional groups tends to be uniform and the density at the surface can be controlled. SAMs provide careful control of specifity of functional group and allow investigation into chain length effects but are model systems rather than applicable systems. SAMs are effective model tools for design and investigation applicable to the optimisation of biomaterials and scaffolds. The importance of protein adsorption from serum must not be forgotten, since this occurs usually prior to cell interactions and differs depending on surface chemistry and topography of the materials (Stanislawski et al., 1995). Protein or other biomolecule adsorption is very easy and provides specific ligands for specific cell interactions. However, it is not suitable for all cell types, mainly those where significant spreading is a disadavantage such as for hepatocytes in liver regeneration.

12.8

Future trends

While surface modification is clearly hugely influencial in controlling the cell response to biomaterials, we must not forget other critical properties that also affect the cell and tissue response. These include matrix stiffness (for more information see publications by Engler et al.) which has been found to influence the differentiation of stem cells, and also the importance of mechanical loading for tissues such as bone and ligament and the benefits of co-culture systems. One final issue to consider is that of 2D versus 3D. 3D culture technologies are on the increase for general cell culture but also for tissue engineering. Significant differences have been observed for several cell types when culture in 2D versus 3D is compared. 3D has all the obvious merits of more closely resembling the in vivo situation but has challenges to overcome in terms of cell viability and mass transfer issues.

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The examples shown here are only a small selection of the wide variety of research in this area. Apologies to those valuable contributors to the field who we could not cite.

12.9

References

Aviss K, Gough JE, Downes S (2010), `Aligned electrospun polymer fibres for skeletal muscle regeneration'. European Cells and Materials 19, 193±204. Bae JW, Lee JH, Choi WS, Lee DS, Bae EH, Park KD (2009), `EPDIM peptideimmobilised porous chitosan beads for enhanced wound healing: Preparation, characterisations and in vitro evaluation', Materials Science and Engineering C 29, 697±701. Beck AJ, Smith-Thomas L, Short RD, MacNeil S (2003), `Development of a plasmapolymerised surface suitable for the transplantation of keratinocyte-melanocyte cocultures for patients with vitiligo', Tissue Engineering 9, 1123±1131. Bush KA, Driscoll, PF, Soto ER, Lambert CR, McGimpsey WG, Pins GD (2008). `Designing tailored biomaterial sufaces to direct keratinocyte morphology, attachment and differentiation', Journal of Biomedical Materials Research Part A 90A, 999±1009. Chen J, Altman GH, Karageorgiou V, Horan R, Collette A, Volloch V, Colabro T, Kaplan DL (2003), `Human bone marrow stromal cell and ligament fibroblast responses on RGD-modified silk fibres', Journal of Biomedical Materials Research Part A 67A, 559±570. Chen X, Wang Z, Qin T-W, Liu C-J, Yang Z-M (2008), `Effects of micropatterned surfaces coated with type I collagen on the proliferation and morphology of tenocytes', Applied Surface Science 255, 368±370. Chiarini A, Petrini P, Bozzini S, Pra ID, Armato U (2003). `Silk fibroin/poly(carbonate)urethane as a substrate for cell growth: in vitro interactions with human cells', Biomaterials 24, 789±799. Cho CS, Seo SJ, Park IK, Kim SH, Kim TH, Hoshiba T, Harada I, Akaike T (2006). `Galactose-carrying polymers as extracellular matrices for liver tissue engineering', Biomaterials 27, 576±585. Clark P, Coles D, Peckham M (1997), `Preferential adhesion to and survival on patterned laminin organizes myogenesis in vitro', Experimental Cell Research 230, 275±283. Clark P, Dunn GA, Knibbs A, Peckham M (2002), `Alignment of myoblasts on ultrafine gratings inhibits fusion in vitro', The International Journal of Biochemistry & Cell Biology 34, 816±825. Cooper E, Parker L, Scotchford CA, Downes S, Leggett GJ, Parker TL (2000), `The effect of alkyl chain length and terminal group chemistry on the attachment and growth of murine 3T3 fibroblasts and primary human osteoblasts on self-assembled monolayers of alkanethiols on gold', Journal of Materials Chemistry 10, 133±139. Dawson RA, Goberdhan NJ, Freedlander E, MacNeil S (1996), `Influence of extracellular matrix proteins on human keratinocyte attachment, proliferation and transfer to a dermal wound model', Burns 22, 93±100. Evans DJR, Britland S, Wigmore PM (1999), `Differential response of fetal and neonatal myoblasts to topographical guidance cues in vitro', Development Genes and Evolution 209, 438±442. Eves PC, Bullett NA, Haddow D, Beck AJ, Layton C, Way L, Shard AG, Gawkrodger DJ, MacNeil S (2008), `Simplifying the delivery of melanocytes and keratinocytes

ß Woodhead Publishing Limited, 2011

Optimising interactions with soft tissues

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for the treatment of vitiligo using a chemically defined carrier dressing', Journal of Investigative Dermatology 128, 1554±1564. Fiegel HC, Havers J, Kneser U, Smith MK, Moeller T, Kluth D, Mooney DJ, Rogiers X, Kaufmann PM (2004), `Influence of flow conditions and matrix coatings on growth and differentiation of three-dimensionally cultured rat hepatocytes', Tissue Engineering 10, 165±174. Fissell WH, Manley S, Westover A, Humes HD, Fleischman AJ, Shuvo R (2006), `Differentiated growth of human renal tubule cells on thin-film and nanostructured materials', ASAIO Journal 52, 221±227. France RM, Short RD, Dawson RA, MacNeil S (1998), `Attachment of human keratinocytes to plasma co-polymers of acrylic acid/octa-1,7-diene an allyl amine/ octa-1,7-diene', Journal of Materials Chemistry 8, 37±42. Garkhal K, Verma S, Tikoo K, Kumar N (2007), `Surface modified poly(L-lactide-co-ecaprolactone) microspheres as scaffold for tissue engineering', Journal of Biomedical Materials Research 82A, 747±756. Gutsche AT , Lo H, Zurlo J, Yager J, Leong KW (1996), `Engineering of a sugarderivatized porous network for hepatocyte culture', Biomaterials 17, 387±393. Haddow DB, Steele DA, Short RD, Dawson RA, MacNeil S (2003), `Plasma polymerised surfaces for culture of human keratinocytes and transfer of cells to an in vitro wound bed model', Journal of Biomedical Materials Research Part A 64A, 80±87. Hodgkinson CP, Wright MC, Paine AJ (2000), `Fibronectin-mediated hepatocyte shape change reprograms cytochrome P450 2C11 gene expression via an integrin-signaled induction of ribonuclease activity', Molecular Pharmacology 58, 976±981. Isaeva VV (1980), `Contact orientation of myoblasts and muscle fibers in the differentiating culture of myogenic cells', Ontogenez 11, 181±187. Kalaskar DM, Gough JE, Ulijn RV, Sampson WW, Scurr DJ, Rutten FJ, Alexander MR, Merry CLR, Eicchorn SJ (2008), `Controlling cell morphology on amino acidmodified cellulose', Soft Matter 4, 1059±1065. Lam MT, Sim S, Zhu X, Takayama S (2006), `The effect of continuous wavy micropatterns on silicone substrates on the alignment of skeletal muscle myoblasts and myotubes', Biomaterials 27, 4340±4347. Lam MT, Huanga Y-C, Birlab RK, Takayama S (2009), `Microfeature guided skeletal muscle tissue engineering for highly organized 3-dimensional free-standing constructs', Biomaterials 30, 1150±1155. Leclerc E, Duval J L, Pezron I, Nadaud F (2008), `Behaviors of liver and kidney explants from chicken embryos inside plasma treated PDMS microchannels', Materials Science and Engineering C 29, 861±868. LeCluyse EL, Bullock PL, Parkinson A (1996), `Strategies for restoration and maintenance of normal hepatic structure and function in long-term cultures of rat hepatocytes', Advanced Drug Delivery Reviews 22, 133±186. Lu HH, Cooper Jr JA, Manuel S, Freeman JW, Attawia MA, Ko FK, Laurencin CT (2005), `Anterior cruciate ligament regeneration using braided biodegradable scaffolds: in vitro optimization studies', Biomaterials 26, 4805±4816. Masuda R, Mochizuki M, Hozumi K, Takeda A, Uchinuma E, Yamashina S, Nomizu M, Kadoya Y (2009), `A novel cell-adhesive scaffold material for delivering keratinocytes reduces granulation tissue in dermal wounds', Wound Repair and Regeneration 17, 127±135. Minoura N, Aiba S, Higuchi M, Gotoh Y, Tsukada M, Imai Y (1995), `Attachment and growth of fibroblast cells on silk fibroin', Biochemical and Biophysical Research Communications 208, 511±516.

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Surface modification of biomaterials

Mochizuki M, Kadoya Y, Wakabayashi Y, Kato K, Okazaki I, Yamada M, Sato T, Sakairi N, Nishi N, Nomizu M (2003), `Laminin-1 peptide-conjugated chitosan membranes as a novel approach for cell engineering', The FASEB Journal 17, 875± 890. Moustafa M, Simpson C, Glover M, Dawson RA, Tesfaye S, Creagh FM, Haddow D, Short R, MacNeil S (2004), `A new autologous keratinocyte dressing treatment for non-healing diabetic neuropathic foot ulcers', Diabetic Medicine 21, 786±798. Sato Y, Terashima M, Kagiwada N, Aung T, Inagaki M, Kakuta T, Saito A (2005), `Evaluation of proliferation and functional differentiation of LLC-PK1 cells on porous polymer membranes for the development of a bioartificial renal tubule device', Tissue Engineering 11, 1506±1515. Seo S-J, Kim I-Y, Choi Y-J, Akaike T, Cho C-S (2006), `Enhanced liver functions of hepatocytes cocultured with NIH 3T3 in the alginate/galactosylated chitosan scaffold', Biomaterials 27, 1487±1495. Stanislawski L, De Nechaud B, Christel P (1995), `Plasma protein adsorption to artificial ligament fibers', Journal of Biomedical Materials Research 29, 315±323. Stern MM, Myers RL, Hammam N, Stern KA, Eberli D, Kritchevsky SB, Soker S, Van Dyke M (2009), `The influence of extracellular matrix derived from skeletal muscle tissue on the proliferation and differentiation of myogenic progenitor cells ex vivo', Biomaterials 30, 2393±2399. Sudhakaran PR, Stamatoglou SC, Hughes RC (1986), `Modulation of protein synthesis and secretion by substratum in primary cultures of rat hepatocytes', Experimental Cell Research 167, 505±516. Tobe S, Takei Y, Kobayashi K, Akaike T (1992), `Tissue reconstruction in primary cultured rat hepatocytes on asialoglycoprotein model polymer', Artificial Organs 16, 526±532. Turner DC, Lawton J, Dollenmeier P, Ehrismann R, Chiquet M (1983), `Guidance of myogenic cell migration by oriented deposits of fibronectin', Developmental Biology 95, 497±504. Wang Y, Kim H-J, Vunjak-Novakovic G, Kaplan DL (2006), `Stem cell-based tissue engineering with silk biomaterials', Biomaterials 27, 6064±6082. Weigel PH (1980), `Rat hepatocytes bind to synthetic galactoside surfaces via a patch of asialoglycoprotein receptors', Journal of Cell Biology 87, 855±861. Weigel PH, Schmell E, Lee YC, Roseman S (1978), `Specific adhesion of rat hepatocytes to -galactosides linked to polyacrylamide gels', Journal of Biological Chemistry 253, 330±333.

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Plate VII Photographs showing de-epidermalized dermis (DED) staining resulting from keratinocyte transfer on (a) (i) hydrocarbon plasma polymer negative control, (ii) uncoated carrier polymer, (iii) rat-tail collagen I positive control. (b) Results of keratinocyte transfer from acrylic acid plasma polymers. Reprinted from Haddow et al. (2003), `Plasma polymerized surfaces for culture of human keratinocytes and transfer of cells to an in vitro wound-bed model', Journal of Biomedical Materials Research, with permission from Wiley.

13

Modifying biomaterial surfaces for the repair and regeneration of nerve cells M . A . M A T E O S - T I M O N E D A and J . A . P L A N E L L , The Institute for Bioengineering of Catalonia (IBEC), Spain and E . E N G E L , Technical University of Catalonia, Spain

Abstract: This chapter discusses the use of chemical and topographical modification of biomaterials in the field of repair and regeneration of nerve tissue. The chapter first reviews briefly the methods developed to chemically and topographically modify the surface of biomaterials. The chapter then discusses how these modified surfaces interact with nerve cells and guide cellular activities. Key words: chemical modification, topographical modification, nerve repair.

13.1

Introduction: the nervous system

The nervous system consists of the central nervous system (CNS) and the peripheral nervous system (PNS). In this chapter, only the CNS will be reviewed in the interests of concision and simplicity.

13.1.1 Central nervous system: cell heterogeneity There are different types of cells in the CNS, including neurons and glial cells. Glial cells comprise astrocytes, oligodendrocytes and microglia. Among these, astrocytes are the predominant glial cell type in the adult mammalian CNS (Richardson, 1980), and together with all other types of glial cells; they outnumber neurons by approximately 10 to 1 (O'Kusky, 1982; White, 2008). The principal characteristics defining mature astrocytes are: a stellate morphology, a key role in the formation of a basal lamina around blood vessels and meninges (Gotow, 1988), and a high expression of intermediate filaments composed of glial fibrilary acidic protein (GFAP) (Eng, 1971; Vaccarino, 2007). Contemporary understanding of astrocytes continues to expand with the discovery of molecular markers which may be used to follow cell lineages in vitro and in vivo. According to current knowledge, astrocytes include a heterogeneous population of cells with diverse morphological features of protein expression,

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varying widely with their development state and distribution in the CNS (Bachoo, 2004; White, 2008). For instance, during development they exist as progenitor cells, called radial glia, while in adulthood. mature astrocytes are subclassified by their morphology and location, as either protoplasmic or fibrous astrocytes. Fibrous astrocytes have fewer, thinner and longer processes than protoplasmic astrocytes. Considerable efforts have been made to reach a clearer definition of subpopulations of astrocytes and their functions based upon the patterns of expression of cellular markers during development, in normal mature nervous system functioning and in pathological conditions (after spinal cord injury, for instance). Neurons are the cells which propagate the signal along the NS. The dogma that these cells could not be replaced throughout human life has prevented advances in neuron tissue regeneration. However, since neural stem cells were discovered (Reynolds, 1992), and were proved to have the ability to regenerate functional neuronal cells, the expectation of treating diseases related to brain ageing or traumas by means of nervous tissue engineering has been greatly raised. These cells have been successfully differentiated into neurons, astrocytes and oligodendrocytes (Reynolds, 1992; McKay, 1997; Cameron, 1998).

13.1.2 Characteristics and general behaviour after injury There are several causes of possible damage to the CNS, such as tumours, vascular diseases or infectious diseases. However, the most common reasons for injury are traumas which are classified as traumatic brain injury (TBI) or traumatic spinal cord injury (SCI). It is estimated that there are between 175 and 200 cases of TBI per 100 000 of the population and 500 cases of SCI per million annually. Most of the cases are the result of traffic accidents and currently there are no effective therapies to heal the damage (data obtained from http://www.guttmann.com). After trauma, SCI results in immediate cell death in the vicinity of the injury site, including that of neurons, glial cells and endothelial cells. The loss of neurons and axon degeneration may cause functional impairment, paraplegia or tetraplegia. The primary injury is followed by a secondary injury cascade including inflammation, degeneration and the formation of an astroglial scar. The scar mainly consists of reactive astrocytes and invading meningeal fibroblasts, which express various inhibitory molecules such as condrotine sulphate proteoglycans. In addition, the injured spinal cord contains inhibitory molecules from myelin debris, such as Nogo and myelin-associated proteoglycan. After TBI, a variety of events take place in the minutes following the injury, which include alterations in cerebral blood flow and the pressure within the skull, both of which contribute substantially to damage from the initial injury.

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After this, the injury site becomes surrounded by a dense glial scar. Reactive astrocytes, glycosaminoglycans and other inhibitory molecules prevent neurons and other cells from infiltrating the injury site, resulting in a loss of axonal connections and a loss of motor function. TBI also causes a cascade of secondary biochemical and molecular changes resulting in delayed tissue damage and cell death, and ultimately to the formation of an irregularly shaped lesion cavity enclosed by a glial scar. The glial scar presents an inhibitory cellular and molecular micro-environment for axon regeneration and tissue repair (Schmidt, 2003).

13.1.3 Astrocytes in injury After a traumatic injury, most of the astrocytes at the lesion site become reactive and form a glial scar. This reaction is induced by environmental cues associated with cell damage and neuro-inflammation and it causes astrocytes to undergo hypertrophy, proliferation, migration, and differentiation to form a dense network bordering the lesion site (White, 2008). Reactive astrocytes increase cell body size, GFAP and GLAST expression (Wu, 1998), and re-express some markers found during the development as Nestin and Vimentin (White, 2008). In addition, they increase the expression of inhibitory extracellular matrix molecules such as chondroitin sulphate proteoglycans, which inhibit neural growth into the lesion (Fitch, 2008). Finally, astrocytes produce chemokines which attract macrophages from the periphery to the site of injury (Otto, 2002; Strack, 2002). However, some studies reported that mature astrocytes are able to de-differentiate into progenitor cells (Lang, 2004) after a CNS injury and to serve as migratory scaffolds for newly generated or transplanted neurons (Harel, 2006), and that radial glial cells are reactivated and up-regulated in areas of neuro-degeneration following spinal cord injury (SCI), thus supplying progenitors for repair (Shibuya, 2003). These recent findings encourage the hypothesis of this study: that glial cells may be directed toward a more permissive state by presenting them with the correct information.

13.1.4 Current therapies Current clinical treatments focus on prevention of secondary injury after the initial trauma and on the partial recovery of function through drug therapy and physical rehabilitation. However, the results are far from ideal and new therapy methods must be established if treatments are to be improved. Systemic delivery of neuro-protective molecules, such as methylprednisolone and gacyclidine, is the only clinically used therapy. While this method has been shown to reduce SCI and to improve functional recovery, it can cause adverse systemic side effects such as immuno-suppression, gastrointestinal bleeding and myopathy (Zhong, 2008).

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Other strategies for the treatment of injuries in the central nervous system are: firstly, the use of artificial electrical prosthesis to recover functionality, a solution already used in some clinical cases. However, it is extremely expensive and not free from drawbacks (Giszter, 2008), and secondly, therapy to counteract the unfavourable glial scar environment. This strategy consists of targeting key molecules involved in the inhibition of the glial scar and therefore of impaired axonal growth. Nogo-A, MAG (myelin associated glicoproteins), Omgp (oligodendrocytes-myelin gliacoprotein) and CSPG (Inhibitory chondroitin sulfate proteoglycans) are some of the molecules identified as inhibiting axon growth, and since their discovery, some neutralising methods have been developed. In the field of neuronal tissue engineering, many researchers are currently focusing their efforts on creating physical or chemical pathways for regenerating axons. These devices include physical or chemical guidance cues, cellular components, and bio-molecular signals, as reviewed by Schmidt and Leach (2003).

13.1.5 Materials for neural regeneration: interaction with cells In response to an injury, cells follow either the route of the compensatory hyperplasia, activating existing adult stem cells, or produce stem cells by the dedifferentiation of mature cells. In all these mechanisms, the regenerationcompetent cells of adult tissues receive information from both physical and chemical cues. This is because in vivo cells exist in three-dimensional environmental `niches' consisting of specific combinations and concentrations of soluble factors and extracellular matrix (ECM) molecules which promote their survival, precisely regulating their proliferation, and determining the phenotypic direction and differentiation. Physical contacts take place in both ECM-cell and material-cell interactions and are regulated by integrin receptors which influence cell migration and growth. Chemical signalling comprises chemo-attractant and chemo-repellents, like netrins, slit proteins or growth factors that are required to direct migrating neural progenitor cells (NPC) to their final destinations (Wyatt Potter, 2008). The goal of bio-material science is to produce supports that can provide a similar level of signalling to ECM. A wide variety of bio-materials are used in the central nervous system for the production of scaffold for tissue regeneration, including both synthetic and natural materials. A schematic list is shown in Table 13.1. Ideally, suitable bio-materials should be easy to make and modify; should exhibit controlled bio-degradation, have good cyto-compatibility, be non-immunogenic, be capable of efficient production and purification, be chemically stable in vivo and have the capacity to enhance specific cell-material interaction (Subramanian, 2009). Most of these materials (particularly the synthetics) have no specific signals which could guide cellular adhesion, orientation, migration or differentiation.

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Table 13.1 Biomaterials used in nerve regeneration Polymer

Description

PLA, PGA and copolymers PLGA

Cell scaffold for transplantion, nerve guidance channel (Wong, 2007; Sundback, 2003) Electrospun cell scaffold (Schnell, 2007; Chew, 2008) Hydrogel with physical characteristics similar to spinal cord (Flynn, 2003; Dalton, 2002) Axonal guidance on nanopatterns or microcontran printed stripes of laminin (Schmalenberg, 2004; Johansson, 2006) Biological function (neurite outgrowth) (Duan, 2007) Peripheral nerve tubulisation (Valero-Cabre, 2004) Cell scaffold, attachment of peptides/polymers (Freudenberg, 2009) Support nerve cell growth (Freier, 2005a, 2005b)

PCL PHEMA PMMA Collagen Silicon PEG Chitosan

Current studies aim to modify the surface properties of these polymers to obtain adequate cues for signalling the cells.

13.2

Surface properties and their effect on neural cells

The new generation of bio-materials display a set of traits which serve as cell instructive materials by directing interactions at the tissue-material interface. They allow a greater control over cell fate and ultimate tissue structure and function. Different materials can induce high variations in cell response because of their inner properties. Surface properties are the main key issues in cellmaterials interaction. Surface properties such as wettability (Soria, 2007), surface chemistry and stiffness (Engler, 2006; Even-Ram, 2006; Georges, 2006) have an influence on cell attachment, viability and differentiation (Hasirci, 2006). Understanding the effects of such surface parameters on cells can help in the design of substrates customised for various uses, e.g. in vitro conditions, to promote or restrain differentiation and conditions that induce selective cell attachment (Wyatt Potter, 2008). Recent studies shown that many cell types respond to micro and nano topographies as they mimic some of the properties of ECM (MartõÂnez, 2009). The development of nanotechnology offers useful tools in designing such a system. The possibility of modifying surfaces, both chemically and physically,

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13.1 Confocal image of a co-culture of glial cells and neurons over PMMA topographically modified with lines of 10 m width.

permits the manufacturing of platforms to control regeneration-competent cell behaviour, either for cell transplant, ex vivo cultures on substrates, or for in situ action (Silva, 2006). More particularly, it has been shown that NPC, Schwann cells and astrocytes preferentially orient themselves along the axis of surface textures. Schwann cells align when cultured on micro-grooved chitosan substrates (Hsu, 2007) or on laminin coated patterned PLA films (Miller, 2001), while Schnell et al. (2007) reported that cells will form dorsal root ganglions (DRG) and olfactory bulb ensheeting cells (OEC) alignment when cultured on electro-spun fibres of meshed poly--caprolactone (PCL)-collagen. PMMA polymers have been shown to promote glial cell attachment while preventing neuron cell attachment (Fig. 13.1). The rigidity of the substrate must be one of the reasons, as the glial cells prefer more rigid surfaces than do neurons. However, neurons attach well on chitosan while glial cells do not. Even though these two surfaces are completely different, it seems that the different chemistry is not the sole cause of their disparate behaviour (Mattoti, 2009).

13.2.1 Surface modifications: technologies and general applications Many methods have been used for physical (topographical features) and chemical modification of the surfaces of both natural and synthetic materials. In

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particular, in the field of neuroscience this entails specific interactions with neurons and glial cells (Nomura, 2006). Physical modifications of surface charge, hydrophilicity and topology have been employed to regulate cell functions, such as attachment, alignment and upregulated protein synthesis (Curtis, 1997). In addition, surface-patterning techniques enable visualisation of the effect of surface properties on cell functions and spatial control of the cellular micro-organisation (Ito, 1999). Many methods have been developed to fabricate precise topography, such as photo-lithography and soft lithography (Curtis, 1998; Folch, 2000). Photolithography consists of the patterning of a layer of photosensitive polymer (photoresist) by means of light (Ristic, 1994). Electron beam lithography uses an electron beam to write directly onto the resist, without the use of a mask. This method, followed by reactive ion etching, allows the fabrication of precise topography with a resolution of about 0.2 m laterally and vertical resolutions of 5 nm, and 5 nm vertically and laterally, respectively. If these methods are used to fabricate masters, they may be used to prepare multiple copies in various polymers by embossing or casting. Embossing involves replication of a network in a polymer substrate from a stamp (or template) fabricated in silicon or metal. Casting entails the evaporation of the solvent of a polymeric solution over a stamp, thus replicating the pattern of the stamp on the polymeric material. Soft lithography is a family of techniques which share the utilisation of a micro-structured surface made of poly(dimethylsiloxane) (PDMS) to generate a pattern (Xia, 1998, 1999). The chemical modification of the surface of the material may be achieved by several methods, depending on the nature of the material to be modified. The surface can be changed chemically by the introduction of functional groups and/ or by covalent coupling of molecules onto the surface. Another approach is the physical adsorption of the molecules over the surfaces (i.e. dip-coating process) (Watanabe, 2007). Many different methods are available for the introduction of functional groups on a surface. In particular, the most common techniques for surface modification in the field of bio-materials are: wet-chemical methods (such as hydrolysis and aminolysis) (Gao, 1998; Zhu, 2002), ozone treatment (Walzak, 1995), UV-treatment and photo-grafting (Kato, 2003; Deng, 2009), selfassembly (Castner, 2002), and plasma treatment (Desmet, 2009). Chemical modification will alter some physical properties such as wettability, but will also enable the subsequent chemical binding of biological molecules to the surface (i.e., extracellular matrix proteins, such as collagen, laminin and fibronectin, as well as small peptide sequences with biological activity derived from them, such as RGD) (Ruoslahti, 1996). Micro-patterning by means of soft lithography and photo-lithography results in both chemical and physical modification of the surface. These techniques lead

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to substrates with regions of precise geometry which permit cell adhesion, while the other areas remains non-adhesive (Kam, 2001; Falconnet, 2006). In a seminal article by Britland et al. (1996), the combination of topography and chemical patterning was examined. It was found that topography is a stronger factor than chemistry when the two are competing, but when both factors are aligned, they act together.

13.2.2 Topographical and chemical surface modifications The cellular response to topographical features represents a well-established phenomenon, as topology has been implicated in processes such as migration and early development (Edgar, 2004). In the field of nerve regeneration, the most commonly used topographical modification is the presence of micro/nanogrooves on the surface (Mukhatyar, 2009). Miller et al. (2001) studied the alignment of Schwann cells into the grooves of PLA (a biodegradable polymer) FDA-approved commonly used for tissue engineering applications which is considered to have a low inflammatory response (Sundback, 2003) The cells aligned in the direction of the grooves, being most influenced by the width of the groove, while the depth does not play a significant role. A similar study was carried out by Li et al. (2008), where the alignment of PC-12 and chick sympathetic neurites on micro-patterned PLA surfaces was examined. The presence of uni-directional grooves as small as 100 nm in height and 1 m in width allows the neurites to align predominantly along the underlying grooves. Moreover, the length of the neurites from chick sympathetic neurons were longer when compared with un-patterned substrates. Analogous studies have been performed with other polymeric materials, such as chitosan and poly(methyl methacrylate) (PMMA). Hsu et al. (2007) investigated the role of micro-grooves in chitosan (a biocompatible and bioactive polymer to nerve cells) (Cheng, 2003; Yuan, 2004). The presence of microgrooves on the surface of the material resulted in the alignment of Schwann cells and glial cells C6 and the expression of genes related to the production of neurotrophic factors. Moreover, the micro-grooved substrates were rolled into conduits and implanted in vivo, showing that peripheral nerve repair on micro-grooved conduits was better than in smooth conduits. However, the chitosan conduits were very fragile and therefore only PLA conduits were tested. Johansson et al. (2006) investigated the axonal outgrowth on nano-imprinted patterns of PMMA. These patterns, with lateral features of 100 nm or larger, can guide axonal outgrowth: a very important feature for nerve-regenerating scaffolds. It is important to notice that in almost all the studies performed with topographically modified substrates, the surfaces had been coated with a layer of laminin to increase initial cell attachment. However, this layer of laminin does not alter the depth of the grooves on the topographically modified substrates

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(Hsu, 2005). Thus, it can be argued that in these studies the substrates are modified both chemically and topographically. Since the patterned and unpatterned surfaces are laminin-coated, the differences observed in cell behaviour can be attributed to the different topography of the substrates. Few studies have been carried out on the effect of both topography and chemistry on the behaviour of hippo-campal and neuronal cells. Li and Folch (2005) have studied the behaviour of axons of embryonic cortical cells when they are presented with competing growth options, i.e. chemical permissiveness and micro-topography. In the same way, Gomez et al. (2007) investigated the effects of immobilised nerve growth factor (NGF) and surface micro-topography on the polarisation and axon elongation of hippo-campal cells. It was found that topography had a more pronounced effect on polarisation compared to immobilised NGF, while the effect of NGF was negligible when both types of cues were presented simultaneously. On the other hand, the axon elongation was affected by NGF immobilisation (10% increase in axon length) whereas there was no effect for the micro-topography. However, there was a synergistic effect in the axon growth when both stimuli were presented (25% increase). Analogous results were obtained for PC12 cells (Foley, 2005), a cell line widely used to investigate neurite outgrowth (Greene 1976). MartõÂnez-Ramos et al. (2008) studied the influence of the surface chemistry of different polymeric materials on the physical absorption of laminin and in the subsequent differentiation of neural stem cells. The polymeric materials range from hydrophobic (poly(ethyl acrylate) (PEA) and poly(methyl acrylate) (PMA)) to hydrophilic homopolymers (poly (hydroxylethyl acrylate) (PHEA) and chitosan), as well as a family of co-polymers of ethyl acrylate (EA) with hydroxyethyl acryklate (HEA) and of methyl acrylate (MA) with HEA with different feeding ratios of the different monomers. This family of co-polymers have similar chemical compositions but different intermediate degrees of hydrophilicity. The results show that the surface chemistry of the substrates influences cell attachment and neuronal differentiation and that co-polymers based on PEA and PHEA in a narrow composition window are suitable substrates for the promotion of cell attachment and the differentiation of neural stem cells into neurons and glia. The difference observed in cell behaviour could be attributed to the different conformation of laminin when absorbed into the substrates (Rodriguez Hernandez, 2007). The bio-functionalisation, i.e. the covalent attachment of molecules with biological activity on the surface of the materials, with proteins or specific active sequences of these proteins, has been the focus of many studies seeking to enhance neuron adhesion and growth (Dodla, 2008; Levesque, 2006). A lamininderived peptide is the most commonly used chemical modification of the surface of materials used for nerve regeneration. This laminin-derived peptide (isoleucine-lysine-valine-alanine-valine (IKVAV)) is known to promote neurite branching and to direct their growth (Matsuzawa, 1996; Yamada, 2002). Other

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Table 13.2 Peptide sequences used for biomaterial functionalisation for nerve regeneration Peptide sequences

Description

RGD PHSRNG6RGD

many ECM proteins: cellular adhesion, migration fibronectin: promote cellular adhesion, migration beyond RGD laminin: cell adhesion, selective neuronal differentiation laminin: cell adhesion laminin: cell adhesion collagen IV: cell adhesion bone marrow homing protein-1: tight adhesion, neurite spreading bone marrow homing protein-2: tight adhesion, neurite spreading

IKVAV YIGSR PDSGR PRGDSGYRGDS SKPPGGTSS PFSSTKT

peptide sequences used to improve neuron adhesion are YIGSR, GRGDS, and PHSRN (see Table 13.2) (Levesque, 2008; Wyatt Potter, 2008). The immobilisation of IKVAV has been employed to enhance the neuronal interaction and nerve regeneration of several materials, such as fluoropolymers (Tong, 2001), poly-L-lysine (PLL) (Thid, 2007), and peptide amphiphiles (PAA) (Silva, 2004). When specific peptide sequences promoting cell-adhesion are presented in a spatially controlled manner (i.e. microcontact printing), neurite outgrowth can be confined to only those areas (Luo, 2004; Burnham, 2006; Hynd, 2007). Moreover, the guiding of neurites has been achieved using concentrated gradients of these peptides (Dertinger, 2002; Adams, 2005).

13.3

Future trends

13.3.1 3D scaffolds based on chemical cues and defined architectures One of the most promising approaches to the treatment of nervous tissue injury is regenerative medicine, and particularly tissue engineering. This application needs `smart' scaffolds with specific physical and chemical cues as well as highly defined architectures which can mimic the tissue structure. The different technologies for obtaining these matrices are discussed below.

13.3.2 Self assembling molecules During the last 10 years, new methods for producing synthetic extracellular matrices have been reported by several groups. Among these, there are two different approaches that we consider to be the most promising in the field of nervous system regeneration.

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13.3.3 Genetically engineered bio-polymers These bio-polymers are designed by means of molecular biotechnology, using bacteria as factories to synthesise large and complex molecules. This is the case of the so called Elastin-like polymers, designed and produced by the RodrõÂguezCabello group in Spain. These polypeptides are composed of repeating sequences found in mammalian elastic protein. The most outstanding features are their elastomeric properties, high bio-compatibility (the immune system cannot differentiate these polymers from mammalian elastin), and above all, the possibility of adding specific peptides to give a signal to tissue related cells. In the case of nervous tissue, the peptide sequences mentioned above as related to neuronal attachment, which promote neurite branching or direct growth, may be tailored to these molecules to stimulate the repair of nervous tissue (Fig. 13.2). Their `smart' and self-assembling capacity relies upon the switch between the folded and unfolded state of the polymer chain as a consequence of a given signal (temperature, pH, light or redox conditions).

13.3.4 Peptide amphiphile These amphiphilic molecules, designed in the laboratory of Prof. Stupp in Chicago, are customised through the peptide sequence for a specific cell response. The nano-fibres self-assemble in aqueous media and place the bioactive epitopes on their surfaces at van der Waals packing distances. These

13.2 ESEM picture of electrospun fibres of an ELP containing the cell-adhesive sequence RGD.

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nano-fibres bundle to form 3D networks and produce a gel-like solid. The nanofibres have high aspect ratios and high surface areas, 5 to 8 nm in diameter and ranging in length from hundreds of nanometres to a few micrometres (Silva, 2004). They have been designed to mimic the collagen structure-building protein-like structural motifs that incorporate sequences of biological interest. In the case of neuronal applications, the molecular design of the scaffold incorporates the pentapeptide epitope isolucinelysine-valine-alanine-valine (IKVAV), which is found in laminin and is known to promote neurite branching and to direct neurite growth (Silva, 2004). In a study performed by this group, it was demonstrated that the high epitope density on the nano-fibres enhanced a rapid differentiation of cells into neurons, so reducing the number of astrocytes.

13.3.5 Electro-spinning Fabrication of micro and nano-fibres using the electro-spinning process is one of the technologies which has aroused more interest in recent years. The advantages of this technique are the simple and cost-effective fabrication process, which does not require high temperatures, and the possibility of producing fibres from precursors which avoid the use of solvents. Matrices are produced by using an electric field to control the deposition of polymer fibres onto a target substrate. The diameters range from several microns to less than 100 nm. Nano-fibres show a high surface area ratio and a highly interconnected porous architecture, which mimics the ECM fibres, allowing cell colonisation, the appropriate diffusion of nutrients and oxygen and the transfer of waste products and CO2. Several polymers have already been used for electro-spinning in neuron applications, including PLA (Yang, 2005). PCL, chitosan, and PCL/chitosan nano-fibres with average fibre diameters of 630, 450, and 190 nm, were produced by electro-spinning. PCL/chitosan scaffolds showed better rat Schwann cell proliferation after eight days of culture than PCL scaffolds and maintained their characteristic cell morphology, with spreading bipolar elongations to the nano-fibrous substrates (Prabhakaran, 2008). As physical features, Christopherson et al. showed that diameter of the fibres has an effect on cell behaviour. Rat hippocampus-derived adult neural stem cells (rNSCs) were cultured on laminin-coated electro-spun polyethersulfone (PES) fibre meshes having an average fibre diameter of 283  45 nm, 749  153 nm and 1452  312 nm. As the diameter of the fibre decreased, a higher degree of proliferation and cell spreading with a lower degree of cell aggregation were observed, indicating the role of fibre topography in regulating differentiation and proliferation of rNSCs in culture (Christopherson, 2009). If the possibility of chemically modifying the fibres by functionalisation with specific peptides is added to this effect, synthetic polymers can be produced which would replicate the signals of the natural ECM.

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13.3.6 Carbon nano-tubes Carbon nano-tubes (CNTs) have electrical, mechanical and chemical properties which make them one of the most promising materials for neuroscience applications. Recent evidence that they are not harmful to human health makes this nano-material a future candidate for use in nervous tissue regeneration. Both single-walled (SWCN) and multi-walled (MWCN) carbon nano-tubes have been increasingly used as scaffolds for neuronal growth and more recently for neural stem cell growth and differentiation. Several groups have successfully cultured neural cells in both SWCN and MWCN (Mattson, 2000), with cone growth and neurite branching (Hu, 2004). However, some limitations in support neurons were indicated when compared to those gathered from neurons grown on widely accepted permissive substrates (Hu, 2004; Liopo, 2006). Several attempts have been made to overcome this drawback and CNTs have been pre-coated with well known adhesive proteins such as collagen or laminin (Nguyen-Vu et al., 2006). The orientation, as well as the purity of the CNTs are relevant for the viability of neurons (Galvan-Garcia, 2007). The surface charge is also relevant, so chemical modification of CNTs using covalent attachments of functional groups (Hu, 2005) has been used to create positively charged surfaces which promote neurite branching and outgrowth. The negative charges showed a decrease in the attachment and survival of cells grown on these substrates (Liopo, 2006). They are also used in interfaces with neurons, where they can detect neuronal electrical activity and deliver electrical stimulation to these cells. The emerging picture is that carbon nano-tubes do not have obvious adverse effects on mammalian health. Thus in the near future, they could be used in brain±machine interfaces.

13.3.7 Computer aided design scaffolds and solid free-form fabrication (rapid prototyping) Solid free-form fabrication (SFF) refers to the physical modelling of a design using a special class of machine technology. SFF systems quickly produce models and prototype parts from 3-D computer-aided design (CAD) model data, computerised tomography (CT), magnetic resonance imaging (MRI) scan data and also from data created by 3-D digitising systems. This is a highly advanced technique capable of fabricating a structure directly from a computer model and can handle complex features, such as internal walls, porosity gradients, complex channels, and multiple material regions. There is the possibility of printing the biomaterial directly or of fabricating a mould. As the process is based on layer-by-layer printing, it is also feasible to combine different materials in order to obtain complex tissues. When related to nervous tissue and focused particularly on the spinal cord, conduits with internal

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channels corresponding to key spinal cord tracts could be created using these approaches. Different types of channels, hollow fibres, desired wall thickness, incorporation of biomolecules and creation of gradients, make this technique a powerful tool in nervous tissue engineering. However, there are still some limitations related to the high cost and the difficulty of incorporating biological components under some processing conditions (Schmidt, 2003).

13.4

References

Adams DN, Kao EY-C, Hypolite CL, Distefano MD, Hu W-S, Letourneau PC 2005, `Growth cones turn and migrate up an immobilised gradient of the laminin IKVAV peptide', Journal of Neurobiology, 62, 1, 134±47. Bachoo RM, Kim RS, Ligon KL, Maher EA, Brennan C, Billings N, Chan S, Li C, Rowitch DH, Wong WH, DePinho RA 2004, `Molecular diversity of astrocytes with implications for neurological disorders', Proceedings of the National Academy of Sciences USA, 101, 22, 8384±9. Britland S, Perridge C, Denyer M, Morgan H, Curtis A, Wilkinson C 1996, `Morphogenetic guidance cues can interact synergistically and hierarchically in steering nerve cell growth', Experimental Biology Online, 1, 2, 1±15. Burnham MR, Turner JN, Szarowski D, Martin DL 2006, `Biological functionalisation and surface micro-patterning of polyacrylamide hydrogels', Biomaterials, 27, 35, 5883±91. Cameron HA, Hazel TG, McKay RDG 1998, `Regulation of neurogenesis by growth factors and neurotransmitters', Journal of Neurobiology, 36, 2, 287±306. Castner DG, Ratner BD 2002, `Biomedical surface science: foundations to frontiers', Surface Science, 500, 1±3, 28±60. Cheng M, Cao W, Gao Y, Gong Y, Zhao N, Zhang Z 2003, `Studies on nerve cell affinity of biodegradable modified chitosan films', Journal of Biomaterials Science, Polymer Edition, 14, 10, 1155±67. Chew SY, Mi R, Hoke A, Leong KW 2008, `The effect of the alignment of electrospun fibrous scaffolds on Schwann cell maturation', Biomaterials, 29, 6, 653±61. Christopherson GY, Song H, Mao H 2009, `The influence of fiber diameter of electrospun substrates on neural stem cell differentiation and proliferation', Biomaterials, 30, 4, 556±64. Curtis A, Wilkinson C 1997, `Topographical control of cells', Biomaterials, 18, 24, 1573±83. Curtis ASG, Wilkinson CDW 1998, `Reactions of cells to topography', Journal of Biomaterials Science-Polymer Edition, 9, 12, 1313±29. Dalton PD, Flynn L, Shoichet MS 2002, `Manufacture of poly(2-hydroxyethyl methacrylate-co-methyl methacrylate) hydrogel tubes for use as nerve guidance channels', Biomaterials, 23, 18, 3843±51. Deng J, Wang L, Liu L, Yang W 2009, `Developments and new applications of UVinduced surface graft polymerizations', Progress in Polymer Science, 34, 2, 156± 93. Dertinger SK, Jiang X, Li Z, Murthy VN, Whitesides GM 2002, `Gradients of substratebound laminin orient axonal specification of neurons', Proceedings of the National Academy of Sciences USA, 99, 20, 12542±7. Desmet T, Morent R, De Geyter N, Leys C, Schacht E, Dubruel P 2009, `Nonthermal

ß Woodhead Publishing Limited, 2011

Repair and regeneration of nerve cells

339

plasma technology as a versatile strategy for polymeric biomaterials surface modification: a review', Biomacromolecules, 10, 9, 2351±78. Dodla MC, Bellamkonda RV 2008, `Differences between the effect of anisotropic and isotropic laminin and nerve growth factor presenting scaffolds on nerve regeneration across long peripheral nerve gaps', Biomaterials, 29, 1, 33±46. Duan X, McLaughlin C, Griffith M, Sheardown H 2007, `Bio-functionalization of collagen for improved biological response: scaffolds for corneal tissue engineering', Biomaterials, 28, 1, 78±88. Edgar D, Kenny S, Almond S, Murray P 2004, ` Topography, stem cell behaviour, and organogenesis', Pediatric Surgery International, 20, 10, 737±40. Eng LF, Vanderhaeghen JJ, Bignami A, Gerstl G 1971, `An acidic protein isolated from fibrous astrocytes', Brain Research, 28, 2, 351±4. Engler AJ, Sen S, Sweeney HL, Discher DE 2006, `Matrix elasticity directs stem cell lineage specification', Cell, 126, 4, 677±89. Even-Ram S, Artym V, Yamada KM 2006, `Matrix control of stem cell fate', Cell, 126, 4, 645±7. Falconnet D, Csucs G, Grandin HM, Textor M 2006, `Surface engineering approaches to micro-pattern surfaces for cell-based assays', Biomaterials, 27, 16, 3044±63. Fitch MT, Silver J 2008, `CNS injury, glial scars, and inflammation: inhibitory extracellular matrices and regeneration failure', Experimental Neurology, 209, 2, 294±301. Flynn L, Dalton PD, Shoichet MS 2003, `Fiber templating of poly(2-hydroxyethyl methacrylate) for neural tissue engineering', Biomaterials, 24, 23, 4265±72. Folch A, Toner M 2000, `Micro-engineering of cellular interactions', Annual Review of Biomedical Engineering, 2, 227±56. Foley JD, Grunwald EW, Nealey PF, Murphy CJ 2005, `Cooperative modulation of neuritogenesis by PC12 cells by topography and nerve growth factor', Biomaterials, 25, 17, 3639±44. Freier T, Montenegro R, Koh HS, Shoichet MS 2005a, `Chitin-based tubes for tissue engineering in the nervous system', Biomaterials, 26, 22, 4624±32. Freier T, Koh HS, Kazazian K, Shoichet MS 2005b, `Controlling cell adhesion and degradation of chitosan films by N-acetylation', Biomaterials, 26, 29, 5872±8. Freudenberga U, Hermann A, Welzel PB, Stirl K, Schwarz SC, Grimmer M, Zieris A, Panyanuwat W, Zschoche S, Meinhold D, Storch A, Werner C 2009, `A star-PEGheparin hydrogel platform to aid cell replacement therapies for neuro-degenerative diseases', Biomaterials, 30, 28, 5049±60. Galvan-Garcia P, Keefer EW, Yang F, Zhang M, Fang S, Zakhidov AA, Baughman RH, Romero MI 2007, `Robust cell migration and neuronal growth on pristine carbon nanotube sheets and yarns', Journal of Biomaterials Science, Polymer Edition, 18, 10, 1245±61. Gao JM, Niklason L, Langer R 1998, `Surface hydrolysis of poly(glycolic acid) meshes increases the seeding density of vascular smooth muscle cells', Journal of Biomedical Material Research, 42, 3, 339±47. Georges PC, Miller WJ, Meaney DF, Sawyer ES, Janmey PA 2006, `Matrices with compliance comparable to that of brain tissue select neuronal over glial growth in mixed cortical cultures', Biophysical Journal, 90, 8, 3012±18. Giszter SF 2008, `Spinal cord injury: present and future therapeutic devices and prostheses', Neurotherapeutics, 5, 1, 147±62. Gomez N, Lu Y, Chen S, Schmidt CE 2007, `Immobilized nerve growth factor and microtopography have distinct effects on polarization versus axon elongation in

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

hippocampal cells in culture', Biomaterials, 28, 2, 271±84. Gotow T, Hashimoto PH 1988, `Deep-etch structure of astrocytes at the superficial glia limitans, with special emphasis on the internal and external organization of their plasma membranes', Journal of Neurocytology, 17, 3, 399±413. Greene LA, Tischler AS 1976, `Establishment of a noradrenergic clonal line of rat adrenal pheochromocytoma cells which respond to nerve growth factor', Proceedings of the National Academy of Sciences USA, 73, 7, 2424±8. Harel NY, Strittmatter SM 2006, `Can regenerating axons recapitulate developmental guidance during recovery from spinal cord injury?', Nature Reviews Neuroscience, 7, 8, 603±16. Hasirci V, Kenar H 2006, `Novel surface patterning approaches for tissue engineering and their effect on cell behavior', Nanomedicine (Lond.), 1, 1, 73±90. Hsu S-H, Chen C-Y, Lu PS, Lai C-S, Chen C-J 2005, `Oriented Schwann cell growth on microgrooved surfaces', Biotechnology and Bioengineering, 92, 5, 579±89. Hsu S, Lu PS, Ni H-C, Su C-H 2007, `Fabrication and evaluation of microgrooved polymers as peripheral nerve conduits', Biomedical Microdevices, 9, 5, 665±74. Hu H, Ni Y, Montana V, Haddon RC, Parpura V 2004, `Chemically functionalized carbon nanotubes as substrates for neuronal growth', Nano Letters, 4, 507±511. Hu H, Ni Y, Mandal SK, Montana V, Zhao B, Haddon RC and Parpura V (2005) `Polyethyleneimine functionalized single-walled carbon nanotubes as a substrate for neuronal growth', Journal of Physical Chemistry B, 109, 4285±9. Hynd MR, Frampton JP, Dowell-Mesfin N, Turner JN, Shain W 2007, `Directed cell growth on protein-functionalized hydrogel surfaces', Journal of Neuroscience Methods, 162, 1±2, 255±63. Ito Y 1999, `Surface micropatterning to regulate cell functions', Biomaterials, 20, 23±24, 2333±42. Johansson F, Carlberg P, Danielsen N, Montelius L, Kanje M 2006, `Axonal outgrowth on nano-imprinted patterns', Biomaterials, 27, 8, 1251±8. Kam L, Shain W, Turner JN, Bizios R 2001, `Axonal outgrowth of hippocampal neurons on micro-scale networks of polylysine-conjugated laminin', Biomaterials, 22, 10, 1049±54. Kato K, Uchide E, Kang ET, Uyama Y, Ikada Y 2003, `Polymer surface with graft chains', Progress in Polymer Science, 28, 2, 209±59. Lang B, Liu HL, Liu R, Feng GD, Jiao XY, Ju G 2004, `Astrocytes in injured adult rat spinal cord may acquire the potential of neural stem cells', Neuroscience, 128, 4, 775±83. Levesque SG, Shoichet MS 2006, `Synthesis of cell-adhesive dextran hydrogels and macroporous scaffolds', Biomaterials, 27, 30, 5277±85. Levesque S, Wylie R, Aizawa Y, Shoichet M 2008, `Peptide modification of polysaccharide scaffolds for targeted cell signaling', in Natural-based polymers for biomedical applications, Woodhead Publishing Ltd, Cambridge. Li J, McNally H, Shi R 2008, `Enhanced neurite alignment on micro-patterned poly-Llactic acid films', Journal of Biomedical Material Research Part A, 87A, 2, 392± 404. Li N, Folch A 2005, `Integration of topographical and biochemical cues by axons during growth on microfabricated 3-D substrates', Experimental Cell Research, 311, 2, 307±16. Liopo AV, Stewart MP, Hudson J, Tour JM, Pappas TC 2006, `Biocompatibility of native and functionalized single-walled carbon nanotubes for neuronal interfaces', Journal of Nanoscience and Nanotechnology, 6, 5, 1365±74.

ß Woodhead Publishing Limited, 2011

Repair and regeneration of nerve cells

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Luo Y, Shoichet MS 2004, `A photolabile hydrogel for guided three-dimensional cell growth and migration', Nature Materials, 3, 4, 249±53. MartõÂnez E, Engel E, Planell JA, Samitier J 2009, `Effects of artificial micro- and nanostructured surfaces on cell behaviour', Annals of Anatomy, 191, 1, 126±35. MartõÂnez-Ramos C, Lainez S, Sancho F, GarcõÂa Esparza MA, Planells-Cases R, GarcõÂa Verdugo JM, GoÂmez Ribelles JL, SalmeroÂn SaÂnchez M, MonleoÂn Pradas M, Barcia JA, Soria J 2008, `Differential of postnatal neural stem cells into glia and functional neurons on laminin-coated polymeric substrates', Tissue Engineering, 14, 8, 1365± 75. Matsuzawa M, Weight FF, Potember RS, Liesi P 1996, `Directional neurite outgrowth and axonal differentiation of embryonic hippocampal neurons are promoted by a neurite outgrowth domain of the B2-chain of laminin', International Journal of Developmental Neuroscience, 14, 3, 283±95. Mattotti M, Ortega A, Delgado L, Planell J, AlcaÂntara S, Engel E 2009, `Differencial behaviour of cortical neural cells related to material properties', ESB 2009 Laussane. Mattson MP, Haddon RC, Rao AM 2000, `Molecular functionalization of carbon nanotubes use as substrates for neuronal growth', Journal of Molecular Neuroscience, 14, 175±82. McKay R 1997, `Stem cells in the central nervous system', Science, 276, 5309, 66±71. Miller C, Shanks H, Witt A, Rutkowski G, Mallapragada S 2001, `Oriented Schwann cell growth on micropatterned biodegradable polymer substrates', Biomaterials, 22, 11, 1263±9. Mukhatyar V, Karumbaiah L, Yeh J, Bellamkonda R 2009, `Tissue engineering strategies designed to realize endogenous regenerative potential of peripheral nerves', Advanced Materials, 21, 46, 4670±9. Nguyen-Vu TDB, Chen H, Cassell AM, Andrews R, Meyyappan M, Li J 2006, `Vertically aligned carbon nanofiber arrays: an advance toward electrical-neural interfaces', Small, 2, 1, 89±94. Nomura H, Tator CH, Shoichet MS 2006, `Bioengineered strategies for spinal cord repair', Journal of Neurotrauma, 23, 3/4, 496±507. O'Kusky J, Colonnier M 1982, `Postnatal changes in the number of astrocytes, oligodendrocytes, and microglia in the visual cortex (area 17) of the macaque monkey: astereological analysis in normal and monocularly deprived animals', The Journal of Comparative Neurology, 210, 3, 307±15. Otto VI, Gloor SM, Frentzel S, Gilli U, Ammann E, Hein AE, Folkers G, Trentz O, Kossmann T, Morganti-Kossmann MC 2002, `The production of macrophage inflammatory protein-2 induced by soluble intercellular adhesion molecule-1 in mouse astrocytes is mediated by src tyrosine kinases and p42/44 mitogen-activated protein kinase', Journal of Neurochemistry, 80, 5, 824±34. Prabhakaran MP, Venugopal JR, Chyan TT, Hai LB, Chan CK, Lim AY, Ramakrishna S 2008, `Electrospun biocomposite nanofibrous scaffolds for neural tissue engineering', Tissue Engineering Part A, 14, 11, 1787±97. Reynolds BA, Weiss S 1992, `Generation of neurons and astrocytes from isolated cells of the adult mammalian central nervous system', Science, 255, 5052, 1707±10. Richardson PM, McGuinness UM, Aguayo AJ 1980, `Axons from CNS neurons regenerate into PNS grafts', Nature, 284, 5753, 264±5. Ristic L 1994, Sensor Technology and Devices, Artech House, Boston. RodrõÂguez Hernandez JC, SalmeroÂn SaÂnchez M, Soria JM, GoÂmez Ribelles JL, MonleoÂn Pradas M 2007, `Substrate chemistry-dependent conformations of single laminin

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

molecules on polymer surfaces are revealed by the phase signal of atomic force microscopy', Biophysical Journal, 93, 1, 202±7. Ruoslahti E 1996, `RGD and other recognition sequences for integrins', Annual Review of Cell and Developmental Biology, 12, 697±715. Schmalenberg KE, Buettner HM, Uhrich KE 2004, `Microcontact printing of proteins on oxygen plasma-activated poly(methyl methacrylate)', Biomaterials, 25, 10, 1851±7. Schmidt C E, Leach JB 2003, `Neural tissue engineering: strategies for repair and regeneration', Annual Review of Biomedical Engineering, 5, 293±347. Schnell E, Klinkhammer K, Balzer S, Brook G, Klee D, Dalton P, Mey J 2007, `Guidance of glial cell migration and axonal growth on electrospun nanofibers of poly-epsiloncaprolactone and a collagen/poly-epsilon-caprolactone blend', Biomaterials, 28, 19, 3012±25. Shibuya S, Miyamoto O, Itano T, Mori S, Norimatsu H 2003, `Temporal progressive antigen expression in radial glia after contusive spinal cord injury in adult rats', Glia, 42, 2, 172±83. Silva GA, Czeisler C, Niece KL, Beniash E, Harrington DA, Kessler JA, Stupp SI 2004, `Selective differentiation of neural progenitor cells by high-epitope density nanofibers', Science, 303, 5662, 1352±5. Silva GA 2006, `Neuroscience nanotechnology: progress, opportunities and challenges', Nature Reviews Neuroscience, 7, 1, 65±74. Soria JM, Martinez Ramos C, Bahamonde O, GarcõÂa Cruz DM, SalmeroÂn SaÂnchez M, GarcõÂa Esparza MA, Casas C, GuzmaÂn M, Navarro X, GoÂmez Ribelles JL, GarcõÂa Verdugo JM, MonleoÂn Pradas M, Barcia JA 2007, `Influence of the substrate's hydrophilicity on the in vitro Schwann cells viability', Journal of Biomedical Materials Research Part A, 83A, 2, 463±70. Strack A, Asensio VC, Campbell IL, Schluter D, Deckert M 2002, `Chemokines are differentially expressed by astrocytes, microglia and inflammatory leukocytes in Toxoplasma encephalitis and critically regulated by interferon-gamma', Acta Neuropathologica, 103, 5, 458±68. Subramanian A, Krishnan UM, Sethuraman S 2009, `Development of biomaterial scaffold for nerve tissue engineering: biomaterial mediated neural regeneration', Journal of Biomedical Science, 16, 108±18. Sundback C, Hadlock T, Cheney M, Vacanti J 2003, `Manufacture of porous polymer nerve conduits by a novel low-pressure injection molding process', Biomaterials, 24, 5, 819±30. Thid D, Bally M, Holm K, Chessari S, Tosatti S, Textor M, Gold J 2007, `Issues of ligand accessibility and mobility in initial cell attachement', Langmuir, 23, 23, 11693±704. Tong YW, Shoichet MS 2001, `Enhancing the neuronal interaction on fluoropolymer surfaces with mixed peptides or spacer group linkers', Biomaterials, 22, 10, 1029± 34. Vaccarino FM, Fagel DM, Ganat Y, Maragnoli ME, Ment LR, Ohkubo Y, Schwartz ML, Silbereis J, Smith KM 2007, `Astroglial cells in development, regeneration, and repair', Neuroscientist, 13, 2, 173±85. Valero-Cabre A, Tsironis K, Skouras E, Navarro X, Neiss WF 2004, `Peripheral and spinal motor reorganization after nerve injury and repair', Journal of Neurotrauma, 21, 1, 95±108. Walzak MJ, Flynn S, Foerch R, Hill JM, Karbashewski E, Lin A, Strobel M 1995, `UV and ozone treatment of polypropylene and poly(ethylene-terephthalate)', Journal of Adhesion Science and Technology, 9, 9, 1229±48. Watanabe J, Nederberg F, Atthoff B, Bowden T, Hilborn J, Ishihara K 2007,

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`Cytocompatible biointerface on poly(lactic acid) by enrichment with phosphorylcholine groups for cell engineering', Material Science and Engineering C: Biomimetic Supramolecular Systems, 27, 2, 3045±51. White RE, Jakeman LB 2008, `Don't fence me in: harnessing the beneficial roles of astrocytes for spinal cord repair', Restorative Neurology and Neuroscience, 26, 2, 197±214. Wong DY, Hollister SJ, Krebsbach PH, Nosrat C 2007, `Poly(epsilon-caprolactone) and poly (L-lactic-co-glycolic acid) degradable polymer sponges attenuate astrocyte response and lesion growth in acute traumatic brain injury', Tissue Engineering, 13, 10, 2515±23. Wu VW, Schwartz JP 1998, `Cell culture models for reactive gliosis: new perspectives', Journal of Neuroscience Research, 51, 6, 675±81. Wyatt Potter REK, Kao WJ 2008, `Biomimetic material systems for neural progenitor cell-based therapy', Frontiers in Bioscience, 13, 806±21. Xia YN, Whitesides GM 1998, `Soft litography', Annual Reviews of Material Science, 28, 153±84. Xia Y, Rogers JA, Paul KE, Whitesides GM 1999, `Unconventional methods for fabricating and patterning nanostructures', Chemical Reviews, 99, 6, 1823±48. Yamada M, Kadoya Y, Kasai S, Kato K, Mochizuki M, Nishi N, Watanabe N, Kleinman HK, Yamada Y, Nomizu M 2002, `Ile-Lys-Val-Ala-Val (IKVAV)-containing laminin 1 chain peptides form amyloid-like fibrils', FEBS Letters, 530, 1±3, 48± 52. Yang F, Murugan R, Wang S, Ramakrishna S 2005, `Electrospinning of nano/micro scale poly(L-lactic acid) aligned fibers and their potential in neural tissue engineering', Biomaterials, 26, 15, 2603±10. Yuan Y, Zhang P, Yang Y, Wang X, Gu X 2004, `The interaction of Schwann cells with chitosan membranes and fibers in vitro', Biomaterials, 24, 18, 4273±78. Zhong Y, Bellamkonda RV 2008, `Biomaterials for the central nervous system', Journal of the Royal Society Interface, 5, 26, 957±75. Zhu YB, Gao CY, Liu XY, Shen JC 2002, `Surface modification of polycaprolactone membrane via aminolysis and biomacromolecule immobilization for promoting cytocompatibility of human endothelial cells', Biomacromolecules, 3, 6, 1312±19.

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Modifying biomaterial surfaces to control stem cell growth and differentiation K . H . S M I T H , Plataforma de Nanotechnologia, Spain and J . W . H A Y C O C K , University of Sheffield, UK

Abstract: The future of tissue engineering and regenerative medicine depends on the development of methods to successfully expand and differentiate stem cells in a specific and reproducible manner. The fabrication of a scaffold that supports the growth and differentiation of stem cells is highly desirable. This chapter addresses the ways in which cell culture surfaces can be modified in order to direct cell behaviour. Surfaces to aid the expansion of stem cells whilst maintaining multipotency are addressed followed by a discussion of surfaces that can be tailored to direct stem cell differentiation along specific lineages. Finally the use of surface modifications to maintain differentiated cells in their terminal phenotype is covered. Key words: stem cells, differentiation, regenerative medicine, tissue engineering, growth factors.

14.1

Introduction

An organ transplant is sometimes the only option for the treatment of a patient. However, there is an extreme shortage of donor tissue and many people die whilst on a transplant waiting list. In addition to treating life-threatening conditions, there is also a need for tissue replacements such as corneas, cartilage and skin grafts. The field of tissue engineering and regenerative medicine seeks to solve the problem of insufficient donor tissue by producing engineered replacements. In general terms the idea is to produce a scaffold seeded with the patient's own cells ± the composite is initially cultured in vitro before implantation and subsequent incorporation into the body. The use of autologous cells gives many advantages over using allogenic or xenogenic cells, as there is less chance of the body rejecting the new tissue. However, there are also problems associated with autologous cell use. The production of tissue usually requires a large number of cells, meaning that a significant biopsy will need to be taken from the patient. This is often not feasible due to the type of cells needed. In addition the tissue to be replaced may be diseased and so the cells may not be suitable. Any harvesting of cells from the patient may also have detrimental consequences (e.g., the harvesting of Schwann cells requires an intact healthy nerve).

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Adult stems cells therefore form an alternative source ± these have been found in most organs of the body, have the ability to self-renew and also differentiate into various cell types. There are a number of potential sites in the body from which a relatively large number of cells can be harvested without significant discomfort for the patient. These include bone marrow, blood and sub-cutaneous adipose tissue. A major challenge in the translational use of stem cells is the requirement to expand the cell population. This needs to be controlled to ensure that proliferation occurs without the loss of multipotency. However, cells often spontaneously differentiate after just a few passages. Under the right conditions adult stem cells can be induced to differentiate into a limited number of different cell types. Defining these conditions is the subject of a great deal of research, and significant progress has been made in recent years. For example, mesenchymal stem cells can be isolated from bone marrow and be differentiated into osteoblasts, chondrocytes or adipocytes. These cells are therefore suitable for bone and cartilage tissue engineering. Adipose-derived stem cells have recently been directed to differentiate into Schwann cells and so will be a useful source of cells for nerve regeneration (e.g. Fig. 14.1). In order to direct the differentiation of these and other types of stem cells, a mixture of growth factors usually needs to be added to the culture medium that cells are grown in. However, growth factors are costly and therefore provide a potential barrier for the scale-up and translation of a tissue engineered product. Using the properties of a cell culture surface to direct differentiation would therefore be a much more cost-effective method. Embryonic stem (ES) cells are also a potential cell source for engineering tissue. Although not autologous, ES cells can produce a lower immune response

14.1 Confocal micrographs of subcutaneous adipose-derived rat stem cells cultured on tissue culture plastic for 10 days in: (a) normal growth medium and (b) Schwann cell differentiation medium. Cells are fluorescently labelled for the Schwann cell markers S100B and GFAP.

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after implantation compared to mature allogenic cells or tissue. Embryonic stem cells are typically harvested from a 4±5-day-old embryo, termed a blastocyst. In contrast to adult stem cells they have the potential to differentiate into almost all of the different cell types of the body. Human embryonic stem cells (hESCs) are typically cultured on a layer of mouse embryonic fibroblasts called a feeder layer. These cells encourage the hESCs to attach and they also secrete the necessary nutrients to keep them in a pluripotent state. However, this use of animal cells is a significant problem if hES cells are to be used for implantation. Not only is there a risk of pathogen transmission but it has been found, for example, that hESCs can incorporate and express the non-human sialic acid Neu5Gc (Martin, 2005). This is likely to produce an immune response due to the presence of the corresponding antibodies in the human host. A great deal of research is therefore focussed on developing methods of culturing hESCs without this feeder layer. The surface on which most cells are cultured is known to have a major impact on their behaviour ± this is now becoming an important area of research for controlling stem cell behaviour. An ideal surface would promote adhesion and proliferation of stem cells whilst maintaining their pluripotency. This chapter therefore reviews approaches for controlling the growth and differentiation of stem cells by modifying the surface chemistry, biochemistry and topography.

14.2

Surfaces for stem cell expansion

In order to population encourages the cells to

use stem cells for the production of engineered tissue a large cell needs to be established. This requires a culture surface that cell adhesion and proliferation whilst maintaining the potential of differentiate into various cell types.

14.2.1 Modification of surface chemistry There are many ways of modifying surfaces in order to increase cell adhesion and proliferation. The surface chemistry can be altered so that the proteins in the cell culture medium adsorb in such a way that the cell binding motifs are displayed in a conformation with a high affinity for the cell surface receptors. The methods involved in producing these surfaces have been discussed in detail in other chapters of this book. For the expansion of stem cells the maintenance of multipotency is an additional complicating factor. For example, it has been shown that altering the wettability of a polymer surface can affect the rate of proliferation of cells ± murine embryonic stem cells (mESCs) proliferate faster on poly(-hydroxy esters) surfaces treated with potassium hydroxide (to increase hydrophilicity) than on un-treated surfaces. In addition to this, these surfaces maintain the cells in an undifferentiated state, as measured by the positive immunoreactivity of Oct-4. This gene encodes a transcription factor necessary for the maintenance of cell pluripotency (Harrison, 2004).

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The effects of different chemical groups on stem cell behaviour have also been investigated. The amination of a polymer surface can significantly increase the proliferation rate of human umbilical cord haematopoietic stem/progenitor cells (HSPCs) compared to un-modified controls (Chua, 2006). The modified surface also increases the expansion efficiency of CD34+ cells, indicating a lower level of differentiation. In another study glass surfaces were modified with silanes terminated in various functional groups (Curran, 2006). Human mesenchymal stem cells were cultured on these surfaces in basal, osteogenic and chondrogenic medium. Cells on the methyl-terminated silane maintained their multipotency even in the differentiation medium whereas cells on amine, thiol, hydroxyl and carboxyl-terminated silanes underwent differentiation. These results quite clearly demonstrate that simple changes to the surface chemistry can have a significant effect on the adhesion, proliferation and differentiation of stem cells. Plasma polymer deposition has been used to form a gradient of the hydrophilic polymer acrylic acid to investigate stem cell behaviour. It was found that mESCs maintained their pluripotency if their level of spreading was restricted by the adhesivity of the surface. At high acid concentrations the cells were well spread but stained only lightly for alkaline phosphatase, a marker of pluripotency. At intermediate acid concentrations cells were less spread but stained strongly for alkaline phosphatase indicating a lack of differentiation (Wells, 2009).

14.2.2 Modification with proteins Modifying surfaces with more complex species is another approach to expanding stem cell populations. Human embryonic stem cells have been cultured on laminin or the basement membrane mimic, matrigel as a replacement for a murine embryonic fibroblast (MEF) feeder layer (Xu, 2001). In the presence of MEF-conditioned medium these cells were capable of undergoing a large number of population doublings whilst expressing the Oct-4 gene and displaying alkaline phosphatase activity. Fibronectin is another protein that has been investigated (Feng, 2006). This is found in the extracellular matrix surrounding most cell types in the body. It has been shown to aid the proliferation of HPSCs when added to culture medium and is thought to be due to interactions of the integrin 4 1 with specific areas of the protein (Yokota, 1998; Schofield, 1998). Fibronectin when covalently attached to a surface (rather than in solution) has also been found to have an effect ± e.g., the expansion rate of CD34+ cells was higher on fibronectin surfaces compared to an un-modified polymer control. Vitronectin has also been found to encourage self-renewal of a number of hESC lines and found to be mediated in part by the v 5 integrin (Braam, 2008).

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14.2.3 Modification with peptides Another approach for controlling the behaviour of stem cells is to covalently attach cell binding peptides to the material surface. There are a number of advantages to using these over proteins. They are made using solid phase synthesis and so are not animal derived. Proteins are likely to undergo conformational changes if adsorbed or covalently attached to a surface, whereas this is not an issue for the short amino acid sequences. Also, it is often possible for cells to adhere to these surfaces even in the absence of animal-derived serum proteins in the cell culture medium. A number of short peptide sequences have been identified that can replicate some of the properties of their native protein. The most commonly used peptide is the arginine-glycine-aspartic acid (RGD) motif found in a number of proteins including fibronectin and collagens. This peptide is highly effective at increasing cell adhesion and proliferation on surfaces to which it is attached. In one study, surfaces were modified by the attachment of fibronectin peptides containing either the RGD motif or the integrin 4 1 binding sequence EILDVPST (Jiang, 2006). The selective expansion of CD34+ HSPCs was found to be greatest on the surface presenting the second of these motifs, with the RGD sequence only providing a small advantage over tissue culture plastic. These results point to a role for the 4 1 integrin in the expansion of HSPCs due to an increase in the capacity of the cells to adhere to the surface.

14.2.4 Modification with growth factors A number of different growth factors are routinely used when expanding stem cells in vitro. Leukaemia inhibitory factor (LIF) is an important cytokine commonly added to cultures of embryonic stem cells (ESCs) to prevent differentiation. The murine embryonic fibroblasts used as a feeder layer for the ES cells also secrete this factor. They produce it as both a soluble form and an immobilised form contained within the extracellular matrix (Rathjen, 1990). It has therefore been hypothesised that LIF may also function when immobilised onto a cell culture surface. In one study the alkaline phosphatase activity of murine ESCs cultured on a LIF-modified polymer scaffold achieved levels similar to that of cells treated with the growth factor in soluble form (Cetinkaya, 2007). The immobilisation of a growth factor may increase its efficiency by reducing internalisation of the receptor-ligand complex by the cell and therefore preventing a reduction in cytokine concentration. Epidermal growth factor (EGF) has also been investigated in an immobilised form. This growth factor is commonly used as a supplement when expanding cultures of neural stem cells. The multipotency of these cells is maintained when cultured on EGF-modified surfaces. The cells proliferated and maintained a stem cell-like phenotype even after four passages (Nakaji-Hirabyashi, 2007). EGF covalently attached to a polymer substrate has

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also demonstrated a protective effect on MSCs by increasing resistance to FasLmediated cell death (Fan, 2007). In comparable experiments soluble EGF did not have the same effect. This was attributed to the prevention of internalisation of the ligand receptor complex due to the chemical attachment.

14.2.5 Enzyme-free sub-culture In order to expand a stem cell population to get an adequate number for clinical use the cells have to go through a number of passages. Traditionally the enzyme trypsin is used to break down the extracellular matrix molecules that anchor the cells to the culture surface. It is a protease produced by the pancreas that works by cleaving certain amide linkages thus breaking proteins down into smaller peptide chains. However, trypsin is not specific to the extracellular matrix and so can damage cell-surface proteins, often to the extent that they become inactive (Canavan, 2005). These same proteins are commonly used as markers for separating stem cells from a mixed cell population. It is also likely that cellsurface receptors such as integrins will be damaged by the process leading to unknown changes in cell signalling. An additional disadvantage to using trypsin for cell expansion is that commercial trypsin is usually animal-derived with recombinant alternatives being prohibitively expensive for clinical use. It is therefore clear that a different approach to lift cells off a culture surface is necessary to overcome these issues. One such approach is to modify the surface with a thermoresponsive polymer that releases the cells from the surface when it is cooled to below a certain temperature (da Silva, 2007). At 37 ëC, the optimum temperature for mammalian cell culture, the polymer would be in a conformation that would present a moderately hydrophobic surface able to support protein adsorption and therefore cell adhesion. Reducing the temperature to below the lower critical solution temperature (LCST) of the polymer causes a change in the conformation that renders the surface significantly more hydrophilic. The polymer chains associate with water molecules rendering protein adsorption energetically unfavourable. The proteins detach from the surface taking the cells with them. The most widely investigated thermoresponsive polymer is poly(Nisopropylacrylamide) (PNIPAM). This has a LCST of 32 ëC in pure water and has been shown to support cell adhesion at 37 ëC making it an ideal candidate for enzyme-free cell detachment. A number of different cell types have been cultured on PNIPAM-grafted surfaces and successfully detached by merely lowering the temperature to below 32 ëC (Akiyama, 2004, Kushida, 2000, Nakajima, 2001, Okano, 1995). An analogous approach by Hopkins (2009) for avoiding trypsin damage to cells is the use of PNIPAM modified with RGD. Here particulates are able to bind to and lift cells off adherent surfaces above an LCST of 34 ëC, followed by a temperature-dependent release of cells below 34 ëC (Hopkins, 2009).

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Pluronic F127 is a triblock co-polymer of poly(ethylene glycol) and poly(propylene glycol) (PEO-PPO-PEO). It changes from an insoluble form displaying a moderately hydrophobic surface to a soluble, hydrophilic form by lowering the temperature to below 30 ëC. L929 fibroblasts attached weakly to a surface that was grafted with F127 and survived for 7 days under standard culture conditions (37 ëC) (Higuchi, 2005, 2006).They were then caused to detach by cooling with the highest efficiency being at the lowest temperature (4 ëC). This same surface was used to culture haematopoietic stem cells (HSCs). In contrast to cells treated with trypsin, HSCs detached from the thermoresponsive Pluronic surface expressed high levels of the stem cell markers CD34 and CD133. It is clear that modifying a cell culture surface with polymers such as these overcomes the major problems with trypsin use. The surfaces have been shown to support the adhesion and growth of a number of different cell types and so it is likely that they would also be able to sustain the growth of different types of stem cells. In addition to this, it has been demonstrated that the detachment process does not damage cell-surface proteins making this technology extremely useful for both stem cell experimentation and translation to the clinic.

14.3

Surfaces for directing stem cell differentiation

As mentioned previously, embryonic stem cells can differentiate into almost all of the cell types of the body. In vivo the cells in the blastocyst first differentiate to form three germ layers, the ectoderm, endoderm and mesoderm. Cells of the ectoderm can further differentiate into the cells of the nervous system, epidermis, eye and mammary glands. Cells of the endoderm produce epithelial tissue and the cells of the liver, pancreas and intestines. Cells of the mesoderm differentiate into skeletal and connective tissue and the cells of the heart, blood and spleen. Cells harvested from a blastocyst and cultured in vitro also have the ability to follow the same differentiation pathways; however, controlling this is complex as cells are intimately dependent on their environment, or niche. Cellcell or cell-ligand interactions trigger different intracellular signalling pathways that result in changes to the cells' phenotype. Different extracellular ligands will bind to different cell-surface receptors triggering various signalling events. Although there are many approaches to manipulating the differentiation of both embryonic and adult stem cells, most of these focus on reproducing the microenvironment of the stem cell niches that direct differentiation in vivo. Ideally this would involve producing a 3D environment with the same mechanical and surface properties, together with control of soluble and insoluble factors such as cytokines, hormones and matrix proteins. Owing to the highly complex nature of this most current research focuses on one specific element, rather than all elements in combination. The ability to modify the in vitro cell culture surface is an important tool for investigating stem cell differentiation. Well-defined surfaces can be produced

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with varying chemistries and topographies. In addition to this, peptides or growth factors can be attached to a surface with the density of the species being highly controllable. It is also possible to pattern surfaces on the micro- or nanoscale in order to investigate more than one factor at a time.

14.3.1 Modification of surface chemistry The chemistry of a cell culture surface affects the conformation of proteins adsorbed to it from the culture medium. This in turn affects the availability of amino acid sequences for cell-surface integrins to bind to. As different integrinligand complexes trigger the different intracellular signalling pathways that result in the differentiation of cells it follows that the conformation of adsorbed proteins could have a direct effect on the resulting phenotype of the cells. The synthesis of well-defined silane monolayers terminated in different functional groups have been used to investigate MSC differentiation (Curran, 2006). Whilst methyl-terminated monolayers maintained the cells in an undifferentiated state the other functional groups supported either chondrogenesis (hydroxyl and carboxyl) or osteogenesis (amino and thiol) in the presence of differentiation medium. Similar work found that in osteogenic medium hydroxyl and aminoterminated monolayers supported differentiation whereas methyl and carboxyl groups did not (Keselowsky, 2005). For this work fibronectin was pre-adsorbed to the surfaces prior to cell seeding, which may account for the difference in findings between these two studies. The proteins adsorbed from culture medium consist of a mixture that includes fibronectin but also a number of other species. These proteins would present a number of integrin binding sequences and so trigger different signalling pathways possibly resulting in alternative directions of differentiation.

14.3.2 Modification of surface topography In vivo, cells are affected by both micro- and nano-scale features in their surrounding environment. The replication of these features provides another approach to directing stem cell behaviour and differentiation. The wide variety of techniques now available for producing topographically different substrates enables detailed investigation of the effects of physical structures. Micro-topography Topographical features on the micro-scale have been shown many times to affect a number of aspects of cell behaviour in vitro. Cells can be encouraged to adhere, spread, proliferate, align and adopt specific shapes by altering the topography of the surface that they are cultured on (Curtis, 1997). It follows that the behaviour of stem cells could be controlled in the same way. The production

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of micro-grooves on a PDMS surface has been shown to cause the alignment and neuronal differentiation of human MSCs (Kim, 2008). Features 1 m wide had the most significant effect with the larger grooves producing less alignment. Calcium levels in the cells were correlated with the level of differentiation. It was found that long neurite extensions increased calcium influx, which in turn increased levels of neuronal markers. By comparing the modified surfaces with flat PDMS it was clear that the topography directly influenced the length and directionality of the neurites resulting in increased levels of neuronal differentiation. Differentiation of neural stem cells (NSCs) into mature neurons can be controlled by altering the size of a honeycomb network of micro-pores on a polymer surface (Tsuruma, 2008). Small pores maintained a degree of multipotency whilst larger pores (10±15 m) encouraged neuronal development measured by MAP2 immunostaining. Nano-topography The effect of nano-scale topography has also been extensively investigated for its effects on cell behaviour (Dalby, 2003; Jin, 2008; Mendonca, 2008). As with micro-scale topographies, stem cells have been shown to respond to surfaces modified on this smaller length scale. Human MSCs can be encouraged to align by nano-gratings formed on a polymer surface (Yim, 2007). Immunostaining showed the presence of neuronal markers such as MAP2 and -tubulin III even in the absence of the differentiating factor retinoic acid (RA). Cells on a flat control surface were only found to be positive for these factors if RA was included in the culture medium. Interestingly, PCR experiments found that MAP2 gene expression of cells grown on nano-structures in the absence of RA was higher than cells growth on flat control surfaces in the presence of RA. This demonstrates that the effect of topography might be more influential than low concentrations of a soluble differentiating factor. The osteogenesis of hMSCs has been shown to be induced by nano-scale roughness produced on a titanium surface using hydrofluoric acid treatment (Valencia, 2009). The expression of osteogenic markers was investigated by PCR and it was clear that cells cultured on the modified surface had adopted a more osteoblastic phenotype. Levels of markers such as a number of bone morphogenic proteins were significantly higher, especially at longer time points. Other surfaces modified to produce roughness on the nano-scale have demonstrated the ability to promote osteogenesis of bone marrow-derived stem cells. These were produced by either a silica nano-particle coating or by an anodisation process to form pores in an aluminium oxide substrate (Lipski, 2007; Popat, 2007). Nano-topography produced on a poly (methyl methacrylate) surface using nano-imprint lithography has been shown to direct the differentiation of bone marrow stromal cells into osteoblast-like cells (Dalby, 2006). Cell spreading and cytoskeletal organisation were increased on the modified

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surfaces compared to the flat control. This was then followed by increases in osteopontin and osteocalcin production at longer time points. This work was then taken further to look at the effect of feature disorder on osteoblastic differentiation (Dalby, 2007). It was found that both fully ordered and completely random patterns produced less evidence of osteogenesis compared to a pattern of features slightly displaced from the ordered array. It was hypothesised that this result may be due to the length of fibrillar adhesions formed by the cells. The slightly displaced features may have allowed longer structures to be achieved and this in turn encourages differentiation. Another highly significant finding was that the slightly displaced features supported osteogenic differentiation of MSCs in the absence of dexamethasone, a promoter of osteogenesis. A number of genes important for osteoblast function were up-regulated in the cells whereas cells cultured under the same conditions but on a planar surface expressed negligible levels of these osteogenic markers. These results demonstrate the level of differentiation control achievable using topographical surface modifications.

14.3.3 Modification with proteins Whilst altering the surface chemistry or topography of a surface can encourage stem cell differentiation, these have a limited range of signalling pathways that they can affect. The controlled presentation of cell-binding ligands using surfaces modified with proteins provides the opportunity to target specific cellsurface receptors and therefore specific intracellular pathways. A large number of peptide ligands have been identified in the major proteins of the extracellular matrix and there are potentially many more to be found in the future. Most approaches to presenting proteins on surfaces involve simple adsorption; however, there are plenty of ways of covalently attaching them via either their amino or carboxy termini, or side chains. There are also methods to attach the molecules in a site-specific manner in order to ensure that access to the cell binding sites is not restricted by changes in protein conformation due to physical interactions with the underlying surface (Reynolds, 2009, Christman, 2007). Fibronectin is found in high concentrations in the extracellular matrix and is also a component of blood plasma. Owing to its widespread presence throughout the body its structure has been investigated thoroughly. A number of cellbinding amino acid sequences have been identified including the RGD motif, which was the first adhesion peptide to be discovered (Pierschbacher, 1984). The RGD-binding integrin 5 1 is termed the fibronectin receptor, but more than half of the integrins known at present recognise the sequence. A number of collagens also contain an RGD sequence amongst their binding sites, although this is often only available to cells after proteolytic degradation (Heino, 2007). Most collagen-binding integrins have 1 as a common sub-unit. When combined with different -units this gives a number of different integrins able to bind to

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different collagens with varying affinities. Collagen type-I is the major collagen found in bone, whereas type-II is more prevalent in cartilage. The basement membrane of cells consists of mainly collagen type-IV in combination with laminin. This is another protein that contains ligands for cell adhesion molecules. The peptides YIGSR and IKVAV are the most commonly studied for their effects on cells. A non-integrin receptor has been found to bind specifically to the YIGSR sequence of laminin whilst the integrins 1 1 and 6 1 have also been identified to interact with the protein (Malino, 1983; Hinek, 1996; Hall, 1990). Between these proteins, and also the many others that cells come into contact with, there are a great number of potential cell-binding sites. The signalling events triggered by these sites are only beginning to be understood. The development of patterning techniques enables the investigation of stem cellprotein interactions to be simplified to a degree. Arrays of different molecules can be produced on surfaces using methods such as micro-contact printing, photolithography or self-assembly. This produces a high-throughput method to enable a number of proteins to be compared (Flaim, 2008). It also facilitates the analysis of soluble growth factor interactions in combination with the effects of the surfaces. Comparisons of fibronectin, laminin and collagens are the most common investigations carried out with regard to directing stem cell differentiation. Although the results of different studies vary greatly, there are some similarities. Fibronectin would appear to increase the presence of hepatocyte markers in murine embryonic stem cells, and neural markers in neural stem cells when cultured in their respective differentiation medium (Flaim, 2005; Ceriotti, 2009). Collagen type-I also up-regulates neural cell markers, either in combination with fibronectin or alone. However, as with many similar findings, the mechanism involved is not fully understood, but would be expected to involve specific sites within the secondary structure of the protein. In contrast, laminin has been shown to increase neural differentiation of both hESCs and neural stem cells in comparison to fibronectin and collagen type I (Ma, 2008, Martinez-Ramos, 2008). Blocking the 6 and 1 integrin sub-units with antibodies partially disrupted the growth of the hESC-derived neural cells indicating a role for this integrin in the laminin-induced effects. This is in line with other work on differentiated cells and has shown a clear link between cell adhesion to laminin and the 6 1 integrin. Although initially appearing as a general `adhesion effect' on differentiated cells this data demonstrates that specific integrin-ligand interactions can have a more precise effect on the behaviour of stem cells. MatrigelTM is a basement membrane mimic that consists of a number of components including the proteins laminin and collagen type IV as the major constituents. It is therefore not surprising that this substance also has an effect on the differentiation of stem cells. Similar to laminin, MatrigelTM can enhance the

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expansion and differentiation of MSCs into neural cells (Qian, 2004). The observed effect was more significant than the use of laminin alone, but the complex nature of MatrigelTM makes it extremely difficult to elucidate an exact component that is responsible. Collagen type IV has also been investigated independently of MatrigelTM. Whilst supporting cell adhesion, this protein was not seen to have any significant effect on stem cell differentiation when compared to other extracellular matrix proteins (Gronthos, 2001; Flaim, 2005; Nakajima, 2007). Thus it is clear that proteins can have specific effects on stem cell differentiation. However, the exact mechanisms are not fully defined or understood, with few cell-ligand interactions identified to date. It is expected that future studies will therefore focus on the underlying mechanisms where specificity is evident.

14.3.4 Modification with peptides An effective way to identify specific amino acid sequences for directing stem cells to differentiate is to study each peptide independently from the rest of the protein. Covalently attaching a peptide to a surface enables it to be presented to a cell in a defined conformation. The peptide is usually not affected by the surface properties to any great extent, which is a major advantage over using the full length protein (where conformation is highly dependent on the underlying surface). Peptides are typically more stable than proteins in cell culture conditions, and are less likely to undergo proteolytic degradation due to the small number of amide linkages present. Specific peptides from the protein laminin have been shown to enhance neurite outgrowth from a number of different neural cell types (Huber, 1998; Richard, 1996; Tashiro, 1989). It therefore follows that these same peptides could be responsible for the induction of neural differentiation seen for the full length protein (Ma, 2008; Martinez-Ramos, 2008). This hypothesis has not been investigated to any great extent. However, one study has found that immobilised peptides containing the YIGSR and IKVAV sequences enhanced neurite outgrowth from neonatal cerebellum stem cells compared to a plain polymer control (He, 2009). This result suggests that there may be a role for these peptides, along with others, in directing neural differentiation of stem cells. In addition to increasing the adhesion of cells to a surface, the RGD peptide has been shown to support the osteogenic differentiation of bone marrow stromal cells (He, 2008; Yang, 2001, 2005). In differentiation medium the cells showed alkaline phosphatase activity and expressed osteogenic markers such as osteocalcin or collagen type I. The RGD-dependence of the effects was verified, whereby soluble RGD peptide was added to the culture medium and a lack of mineralisation was seen.

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14.3.5 Modification with growth factors In vitro differentiation of both embryonic and adult stem cells involves the addition of a number of different soluble growth factors to the culture medium. Osteogenic differentiation, for example, requires a combination of -glycerophosphate, ascorbic acid and dexamethasone. Although these conditions are now well established, their mechanisms of action are less well understood. It is known that on binding, some of the receptor ligand complexes are internalised by the cell, whereas others remain on the outside. It has also been found that some growth factors normally internalised can function if immobilised onto a surface. This can occur in vivo as well as in vitro if the factor is attached to an extracellular matrix protein. It has been hypothesised previously that stem cells can be directed down a specific differentiation pathway when grown on a surface modified with growth factors. Due to the large number of growth factors that have been identified as having an effect on stem cell differentiation patterned arrays are often used. These generally consist of a cell-adhesion protein co-immobilised with a growth factor. Using this method, the differentiation of neural stem cells (NSCs) into neuronal cells has been enhanced by immobilised nerve growth factor (NGF) and neutrophin-3 (NT-3) whereas ciliary neurotrophic factor (CNTF) promoted an astroglial phenotype (Nakajima, 2007). The astroglial differentiation was notably higher when the CNTF was co-immobilised with either fibronectin or an RGD-containing artificial protein. This shows that immobilised extracellular matrix proteins and growth factors can work synergistically to direct stem cell differentiation. To remove the potential variation in results due to the adhesion protein, other work has used laminin alone to promote the cell attachment. The growth factors were then co-immobilised in an array. Glial differentiation was seen for NSCs cultured on surfaces with immobilised CNTF, transforming growth factor- (TGF- ) and bone morphogenic protein-4 (BMP-4). Specific intracellular signalling mechanisms were also investigated by attaching ligands that induce the notch or wnt pathways. The notch pathway was implicated in the promotion of glial differentiation whilst preventing neuronal differentiation. Coimmobilisation of CNTF and notch ligands such as jagged-1 increased the level of glial differentiation even further. These arrays of growth factors enable the complex nature of stem cell differentiation to be investigated in detail. Even though current understanding is limited, the use of these methods has provided vital information and shows that the use of immobilised growth factors has enormous potential for the future.

14.4

Surfaces for maintaining the differentiated cell phenotype

A variety of different surface modifications have been shown to have a significant impact on the in vitro differentiation of stem cells. However, an

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important question still to be answered is how will the cells behave once implanted into the body? Cells can readily be encouraged down a specific lineage in vitro but the change in environment upon implantation will potentially cause cells to de-differentiate or trans-differentiate. It is also possible that the cells may retain some of their stem cell properties and proliferate uncontrollably causing tumour formation. The presence of soluble growth factors can no longer be regulated and so the scaffold that the cells are implanted on or in becomes even more important in determining the cell phenotype. Tissue engineering scaffolds generally provide a three-dimensional environment in which cells can proliferate and produce their own extracellular matrix. This will then replace the scaffold as it gradually degrades. In vivo, cells exist in a 3D matrix and so it is not surprising that stem cell-derived neurons have been shown to produce more extensive neurite growth and higher expression of neural markers when cultured in 3D in comparison to 2D (Hayman, 2004). The physical properties of the scaffold have also been seen to affect the phenotype of cells. The differentiation pathway of MSCs can be tailored by culturing them on gels with different elastic moduli (Engler, 2006; Rowlands, 2008). Cells cultured on gels with high rigidity are more likely to undergo osteogenesis, with more flexible substrates encouraging the formation of a neural phenotype. Tailoring the physical properties of the scaffold is therefore an important step towards controlling the differentiated cell phenotype. Another level of control over the cells is the surface that they are in contact with. It is clear that modifying the surface chemistry, topography or biomolecule presentation can direct the differentiation of stem cells down specific lineages ± but can these same modifications maintain the differentiated cell phenotype? Modifying the chemistry of a surface alters the type and conformation of adsorbed serum proteins, which in turn determines ligand presentation to cellsurface receptors. Although this is a useful way of controlling stem cell differentiation in vitro, it is unlikely to be as effective in vivo. Some studies have demonstrated levels of stem cell differentiation on surfaces modified with nanotopography even in the absence of neurogenic or osteogenic growth factors (Yim, 2007, Dalby, 2007). These are encouraging results as they show control of cell phenotype can be gained with a simple change in surface topography. The application of the same topographical features to the surface of an implant may enable maintenance of the differentiated phenotype in vivo. Laminin has significant effects on the phenotype of different cell types. In vivo, adult smooth muscle cells (SMCs) have a bi-polar morphology, do not proliferate, do not synthesise extracellular matrix and are able to contract. In contrast, the conditions of in vitro culture change the cells to a more fibroblastlike morphology followed by extensive proliferation and protein synthesis. This phenotype is analogous to that of SMCs found in arteries damaged by atherosclerosis. Here, cells proliferate into the intimal layer of the artery increasing the level of occlusion. The maintenance of a contractile phenotype of

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14.2 Confocal micrographs of NG108-15 neuronal cells cultured on substrates of: (a) fibronectin and (b) laminin for 2 days. Cells are fluorescently labelled for F-actin filaments.

stem-cell derived SMCs is therefore necessary to prevent this highly undesirable response. Freshly harvested SMCs cultured on a laminin-modified substrate have been shown to maintain their contractile phenotype for a longer time than on other proteins such as fibronectin, which actively promotes SMC proliferation (Hayward, 1995; Hirst, 2000; Roy, 2002). Neurite outgrowth from neural cells is increased dramatically by culture on laminin (e.g. Fig. 14.2). Both the length and number of neurites per cell are higher on this substrate compared to un-modified polymer surfaces and other proteins (Carri, 1988; Kleinman, 1988). These results indicate that modifying a surface with laminin could be an effective way to maintain the phenotype of stem cell-derived SMCs or neural cells on implantation into the body. Collagen type IV is another major protein of the basement membrane. It can also delay the transformation of SMCs from a contractile to a synthetic phenotype in culture. As with laminin, collagen IV-modified substrates have been shown to encourage neurite outgrowth from neurons cultured on the surface (Turner, 1989; Lein, 1991). The levels of outgrowth have been shown to be higher than on fibronectin or poly (D-lysine), a commonly used adhesive substrate for neural cells. Other studies have shown that modifying a surface with a combination of laminin and collagen type IV (or with the artificial basement membrane mixture Matrigel) also maintain SMC contractile phenotype and promote neurite outgrowth (Thyberg, 1994, Hirst, 2000). It is therefore clear that mimicking the basement membrane is an effective method to maintain the phenotype of both SMCs and neural cells. It is also likely that many other cell types could be affected in the same way. As mentioned previously, using short peptides to modify a surface has advantages over the use of full-length proteins. These advantages are amplified

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for a substrate that is to be implanted into the body. The presence of other cells and enzymes makes degradation of an immobilised protein much more likely and the introduction of animal-derived compounds to the human body introduces the risk of pathogen transmission. The laminin peptides YIGSR, IKVAV and LQVQLSIR have all demonstrated the ability to promote neurite outgrowth when immobilised onto a cell culture substrate (Huber, 1998; Richard, 1996; Tashiro, 1989). In one study neurite outgrowth was more significant on YIGSR covalently bound to an alginate surface than on laminin adsorbed onto it (Dhoot, 2004). In this case it is possible that the conformation of the adsorbed protein has a reduced neuritepromoting effect. The peptide RGD is found in a number of proteins including fibronectin and fibrillin-1. As with these full-length proteins a surface modified with this peptide transforms freshly isolated contractile smooth muscle cells into their synthetic, proliferative phenotype (Hedin, 1989). Both laminin and collagen type IV contain RGD motifs; however, studies have shown that cell adhesion to these proteins is not dependent on this sequence (Aumailley, 1990; Herbst, 1988). It is likely that it is hidden within the proteins making it unavailable for integrin binding. This suggests that the RGD peptide is specifically involved in the modulation of SMCs from a contractile to a synthetic phenotype.

14.5

Future trends

It is clear that tremendous potential exists in the use of both adult and embryonic stem cells for regenerative medicine and cellular therapies. Whilst the use of culture surfaces to control the expansion and differentiation of these cells is only beginning to be explored in detail, results obtained so far show a lot of promise for this area of research. The commercialisation of stem cell products and therapies will not be possible without a number of hurdles being overcome. The first of these is removing the need for murine feeder layers and other animalderived products. We have seen that modifying the culture surface can go some way to solving this problem. Controlling the surface chemistry or topography and the immobilisation of synthetic peptides can be used to aid the expansion and direct the differentiation of stem cells. Further research into these areas may produce methods that could replace the use of animal products altogether. Further difficulties with the commercialisation of stem cells are scale-up and reproducibility. The use of artificial surfaces, instead of a cell feeder layer plus numerous soluble growth factors, could go a long way to solving these problems as most surface modifications are highly reproducible over large areas. Finally, the cost-effectiveness of cell-based therapies is a presently a barrier to commercialisation. Current methods for expanding and differentiating stem cells involves a large number of soluble growth factors ± these are extremely expensive and are likely make any attempt to market a stem cell product almost

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impossible. The use of modified culture surfaces may moderate production costs by reducing the need for growth factors. Many of the methods employed are relatively cheap and their potential as substrates for stem cell culture is gaining increased recognition. Future research in this area is needed to more fully understand exactly how these surfaces articulate their responses and how they can be modified to obtain higher levels of control over the behaviour of stem cells. The use of immobilised peptides also cost less than that of their soluble counterpart proteins and these ligands, too, can control the specific response of stem cells. At present the understanding of the signalling mechanisms involved is incomplete, but it is likely that further research here will provide a more detailed understanding. Thus the huge number of potential ligands for cell-surface receptors, topographies and physical properties may enable considerable control over stem cell behaviour in the future ± simply by culturing cells on a modified surface.

14.6

Acknowledgments

We are grateful to Rossukon Kaewkhaw and Celia Murray-Dunning (Kroto Research Institute, University of Sheffield, UK) for the data used in Figs 14.1 and 14.2.

14.7

References

Akiyama Y, Kikuchi A, Yamato M, Okano T (2004), `Ultrathin poly(Nisopropylacrylamide) grafted layer on polystyrene surfaces for cell adhesion/ detachment control', Langmuir, 20, 5506±5511. Aumailley M, Gerl M, Sonnenberg A, Deutzmann R, Timpl R (1990), `Identification of the Arg-Gly-Asp sequence in laminin A chain as a latent cell-binding site being exposed in fragment P1', FEBS Lett, 262, 82±86. Braam SR, Zeinstra L, Litjens S, Ward-van Oostwaard D, van den Brink S, van Laake L, Lebrin F, Kats P, Hochstenbach R, Passier R, Sonnenberg A, Mummery CL (2008), `Recombinant vitronectin is a functionally defined substrate that supports human embryonic stem cell self-renewal via alpha V beta 5 integrin', Stem Cells, 26, 2257± 2265. Canavan HE, Cheng X, Graham DJ, Ratner BD, Castner DG (2005), `Cell sheet detachment affects the extracellular matrix: A surface science study comparing thermal liftoff, enzymatic and mechanical methods, J Biomed Mater Res A, 75A, 1±13. Carri NG, Perris R, Johansson S, Ebendal T (1988), `Differential outgrowth of retinal neurites on purified extracellular matrix molecules', J Neurosci Res, 19, 428±439. Ceriotti L, Buzanska L, Rauscher H, Mannelli I, Sirghi L, Gilliland D, Hasiwa M, Bretagnol F, Zychowicz M, Ruiz A, Bremmer S, Coeke S, Colpo P, Rossi F (2009), `Fabrication and characterization of protein arrays for stem cell patterning', Soft Matter, 5, 1406±1416. Cetinkaya G, Turkoglu H, Arat S, Odaman H, Onur MA, Gumusderelioglu M, Tumer A (2007), `LIF-immobilized nonwoven polyester fabrics for cultivation of murine embryonic stem cells', J Biomed Mater Res A, 81A, 911±919.

ß Woodhead Publishing Limited, 2011

Stem cell growth and differentiation

361

Christman KL, Broyer RM, Tolstyka ZP, Maynard HD (2007), `Site-specific protein immobilization through N-terminal oxime linkages', J Mater Chem, 17, 2021±2027. Chua KN, Chai C, Lee PC, Tang YN, Ramakrishna S, Leong KW, Mao HQ (2006), `Surface-aminated electrospun nanofibers enhance adhesion and expansion of human umbilical cord blood hematopoietic stem/progenitor cells', Biomaterials, 27, 6043±6051. Curran JM, Chen R, Hunt JA (2006), `The guidance of human mesenchymal stem cell differentiation in vitro by controlled modifications to the cell substrate', Biomaterials, 27, 4783±4793. Curtis A, Wilkinson C (1997), `Topographical control of cells', Biomaterials, 18, 1573± 1583. Dalby MJ, Childs S, Riehle MO, Johnstone HJH, Affrossman S, Curtis ASG (2003), `Fibroblast reaction to island topography: changes in cytoskeleton and morphology with time', Biomaterials, 24, 927±935. Dalby MJ, McCloy D, Robertson M, Agheli H, Sutherland D, Affrossman S, Oreffo ROC (2006), `Osteoprogenitor response to semi-ordered and random nanotopographies', Biomaterials, 27, 2980±2987. Dalby MJ, Gadegaard N, Tare R, Andar A, Riehle MO, Herzyk P, Wilkinson CDW, Oreffo ROC (2007), `The control of human mesenchymal cell differentiation using nanoscale symmetry and disorder', Nat Mater, 6, 997±1003. da Silva RMP, Mano JF, Reis RL (2007), `Smart thermoresponsive coatings and surfaces for tissue engineering: Switching cell-material boundaries', Trends Biotechnol, 25, 577±583. Dhoot NO, Tobias CA, Fischer I, Wheatley MA (2004), `Peptide-modified alginate surfaces as a growth permissive substrate for neurite outgrowth' J Biomed Mater Res A, 71A, 191±200. Engler AJ, Sen S, Sweeney HL, Discher DE (2006), `Matrix elasticity directs stem cell lineage specification', Cell, 126, 677±689. Fan VH, Au A, Tamama K, Littrell R, Richardson LB, Wright JW, Wells A, Griffith LG (2007), `Tethered epidermal growth factor provides a survival advantage to mesenchymal stem cells', Stem Cells, 25, 1241±1251. Feng Q, Chai C, Jiang XS, Leong KW, Mao HQ (2006), `Expansion of engrafting human hematopoietic stem/progenitor cells in three-dimensional scaffolds with surfaceimmobilized fibronectin', J Biomed Mater Res A, 78A, 781±791. Flaim CJ, Chien S, Bhatia SN (2005), `An extracellular matrix microarray for probing cellular differentiation', Nat Methods, 2, 119±125. Flaim CJ, Teng D, Chien S, Bhatia SN (2008), `Combinatorial signaling microenvironments for studying stem cell fate', Stem Cell Dev, 17, 29±39. Gronthos S, Simmons PJ, Graves SE, Robey PG (2001), `Integrin-mediated interactions between human bone marrow stromal precursor cells and the extracellular matrix', Bone, 28, 174±181. Hall DE, Reichardt LF, Crowley E, Holley B, Moezzi H, Sonnenberg A, Damsky CH (1990), `The alpha-1/beta-1 and alpha-6/beta-1integrin heterodimers mediate cell attachment to distinct sites on laminin', J Cell Biol, 110, 2175±2184. Harrison J, Pattanawong S, Forsythe JS, Gross KA, Nisbet DR, Beh H, Scott TF, Trounson AO, Mollard R (2004), `Colonization and maintenance of murine embryonic stem cells on poly(alpha-hydroxy esters)', Biomaterials, 25, 4963±4970. Hayman MW, Smith KH, Cameron NR, Przyborski SA (2004), `Enhanced neurite outgrowth by human neurons grown on solid three-dimensional scaffolds', Biochem Biophys Res Commun, 314, 483±488.

ß Woodhead Publishing Limited, 2011

362

Surface modification of biomaterials

Hayward IP, Bridle KR, Campbell GR, Underwood PA, Campbell JH (1995), `Effect of extracellular matrix proteins on vascular smooth muscle cell phenotype', Cell Biol Int, 19, 839±846. He LM, Liao SS, Quan DP, Ngiam M, Chan CK, Ramakrishna S, Lu J (2009), `The influence of laminin-derived peptides conjugated to Lys-capped PLLA on neonatal mouse cerebellum C17.2 stem cells', Biomaterials, 30, 1578±1586. He XZ, Ma JY, Jabbari E (2008), `Effect of grafting RGD and BMP-2 protein-derived peptides to a hydrogel substrate on osteogenic differentiation of marrow stromal cells', Langmuir, 24, 12508±12516. Hedin U, Bottger BA, Luthman J, Johansson S, Thyberg J (1989), `A substrate of the cellattachment sequence of fibronectin (Arg-Gly-Asp-Ser) is sufficient to promote transition of arterial smooth muscle cells from a contractile to a synthetic phenotype', Dev Biol, 133, 489±501. Heino J (2007), `The collagen family members as cell adhesion proteins', Bioessays, 29, 1001±1010. Herbst TJ, McCarthy JB, Tsilibary EC, Furcht LT (1988), `Differential effects of laminin, intact type IV collagen and specific domains of type IV collagen on endothelial cell adhesion and migration', J Cell Biol, 106, 1365±1373. Higuchi A, Yamamoto T, Sugiyama K, Hayashi S, Tak TM, Nakagawa T (2005), `Temperature-dependent cell detachment on Pluronic gels, Biomacromolecules, 6, 691±696. Higuchi A, Aoki N, Yamamoto T, Miyazaki T, Fukushima H, Tak TM, Jyujyoji S, Egashira S, Matsuoka Y, Natori SH (2006), `Temperature-induced cell detachment on immobilized pluronic surface', J Biomed Mater Res A, 79A, 380±392. Hinek A (1996), `Biological roles of the non-integrin elastin/laminin receptor', Biol Chem, 377, 471±480. Hirst SJ, Twort CHC, Lee TH (2000), `Differential effects of extracellular matrix proteins on human airway smooth muscle cell proliferation and phenotype', Am J Respir Cell Mol Biol, 23, 335±344. Hopkins S, Carter SR, Haycock JW, Fullwood NJ, MacNeil S, Rimmer S (2009), `Submicron poly(N-isopropyl acrylamide) particles as temperature responsive vehicles for the detachment and delivery of human cells', Soft Matter, 5(24), 4928±4937. Huber M, Heiduschka P, Kienle S, Pavlidis C, Mack J, Walk T, Jung G, Thanos S (1998), `Modification of glassy carbon surfaces with synthetic laminin-derived peptides for nerve cell attachment and neurite growth', J Biomed Mater Res, 41, 278±288. Jiang XS, Chai C, Zhang Y, Zhuo RX, Mao HQ, Leong KW (2006), `Surfaceimmobilization of adhesion peptides on substrate for ex vivo expansion of cryopreserved umbilical cord blood CD34(+) cells', Biomaterials, 27, 2723±2732. Jin CY, Zhu BS, Wang XF, Lu QH, Chen WT, Zhou XJ (2008), `Nanoscale surface topography enhances cell adhesion and gene expression of madine darby canine kidney cells', JMater Sci-Mater Med,19, 2215±2222. Keselowsky BG, Collard DM, Garcia AJ (2005), `Integrin binding specificity regulates biomaterial surface chemistry effects on cell differentiation', Proc Natl Acad Sci USA,102, 5953±5957. Kim SJ, Lee JK, Kim JW, Jung JW, Seo K, Park SB, Roh KH, Lee SR, Hong YH, Lee YS, Kang KS (2008), `Surface modification of polydimethylsiloxane (PDMS) induced proliferation and neural-like cells differentiation of umbilical cord bloodderived mesenchymal stem cells', J Mater Sci-Mater Med, 19, 2953±2962. Kleinman HK, Ogle RC, Little CD, Sweeney TM, Luckenbilledds L (1988), `Laminin receptors for neurite formation', Proc Natl Acad Sci USA, 85, 1282±1286.

ß Woodhead Publishing Limited, 2011

Stem cell growth and differentiation

363

Kushida A, Yamato M, Konno C, Kikuchi A, Sakurai Y, Okano T (2000), `Temperatureresponsive culture dishes allow nonenzymatic harvest of differentiated MadinDarby Canine Kidney (MDCK) cell sheets', J Biomed Mater Res, 51, 216±223. Lein PJ, Higgins D, Turner DC, Flier LA, Terranova VP (1991), `The NC1 domain of type IV collagen promotes axonal growth in sympathetic neurons through interaction with the alpha-1 beta-1 integrin', J Cell Biol, 113, 417±428. Lipski AM, Jaquiery C, Choi H, Eberli D, Stevens M, Martin I, Chen IW, Shastri VP (2007), `Nanoscale engineering of biomaterial surfaces', Adv Mater, 19, 553±557. Ma W, Tavakoli T, Derby E, Serebryakova Y, Rao MS, Mattson MP (2008), `Cellextracellular matrix interactions regulate neural differentiation of human embryonic stem cells', BMC Dev Biol, 8, 90. Malinoff HL, Wicha MS (1983), `Isolation of a cell surface receptor protein for laminin from murine fibrosarcoma cells', J Cell Biol, 96, 1475±1479. Martin MJ, Muotri A, Gage F, Varki A (2005), `Human embryonic stem cells express an immunogenic nonhuman sialic acid', Nat Med, 11, 228±232. Martinez-Ramos C, Lainez S, Sancho F, Esparza MAG, Planells-Cases R, Verdugo JMG, Ribelles JLG, Sanchez MS, Pradas MM, Barcia JA, Soria JA (2008), `Differentiation of postnatal neural stem cells into glia and functional neurons on laminin-coated polymeric substrates', Tiss Eng A,14, 1365±1375. Mendonca G, Mendonca DBS, Aragao FJL, Cooper LF (2008), `Advancing dental implant surface technology ± from micron- to nanotopography', Biomaterials, 29, 3822±3835. Nakaji-Hirabayashi T, Kato K, Arima Y, Iwata H (2007), `Oriented immobilization of epidermal growth factor onto culture substrates for the selective expansion of neural stem cells', Biomaterials, 28, 3517±3529. Nakajima K, Honda S, Nakamura Y, Lopez-Redondo F, Kohsaka S, Yamato M, Kikuchi A, Okano T (2001), `Intact microglia are cultured and non-invasively harvested without pathological activation using a novel cultured cell recovery method', Biomaterials, 22, 1213±1223. Nakajima M, Ishimuro T, Kato K, Ko IK, Hirata I, Arima Y, Iwata H (2007), `Combinatorial protein display for the cell-based screening of biomaterials that direct neural stem cell differentiation', Biomaterials, 28, 1048±1060. Okano T, Yamada N, Okuhara M, Sakai H, Sakurai Y (1995), `Mechanism of cell detachment from temperature-modulated, hydrophilic±hydrophobic polymer surfaces', Biomaterials, 16, 297±303. Pierschbacher MD, Ruoslahti E (1984), `Cell attachment activity of fibronectin can be duplicated by small synthetic fragments of the molecule', Nature, 309, 30±33. Popat KC, Chatvanichkul KI, Barnes GL, Latempa TJ, Grimes CA, Desai TA (2007), `Osteogenic differentiation of marrow stromal cells cultured on nanoporous alumina surfaces', J Biomed Mater Res A, 80A, 955±964. Qian LC, Saltzman WM (2004), `Improving the expansion and neuronal differentiation of mesenchymal stem cells through culture surface modification', Biomaterials, 25, 1331±1337. Rathjen PD, Toth S, Willis A, Heath JK, Smith AG (1990), `Differentiating activity is produced in matrix-associated and diffusible forms that are generated by alternate promoter usage', Cell, 62, 1105±1114. Reynolds NP, Tucker JD, Davison PA, Timney JA, Hunter CN, Leggett GJ (2009), `Sitespecific immobilization and micrometer and nanometer scale photopatterning of yellow fluorescent protein on glass surfaces', J Am Chem Soc, 131, 896±897. Richard BL, Nomizu M, Yamada Y, Kleinman HK (1996), `Identification of synthetic

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

peptides derived from laminin alpha 1 and alpha 2 chains with cell type specificity for neurite outgrowth', Exp Cell Res, 228, 98±105. Rowlands AS, George PA, Cooper-White JJ (2008), `Directing osteogenic and myogenic differentiation of MSCs: interplay of stiffness and adhesive ligand presentation', Am J Physiol Cell Physiol, 295, C1037±C1044. Roy J, Tran PK, Religa P, Kazi M, Henderson B, Lundmark K, Hedin U (2002), `Fibronectin promotes cell cycle entry in smooth muscle cells in primary culture', Exp Cell Res, 273, 169±177. Schofield KP, Humphries MJ, de Wynter E, Testa N, Gallagher JT (1998), `The effect of alpha(4)beta(1)-integrin binding sequences of fibronectin on growth of cells from human hematopoietic progenitors', Blood, 91, 3230±3238. Tashiro K, Sephel GC, Weeks B, Sasaki M, Martin GR, Kleinman HK, Yamada Y (1989), `A synthetic peptide containing the IKVAV sequence from the A-chain of laminin mediates cell attachment, migration and neurite outgrowth', J Biol Chem, 264, 16174±16182. Thyberg J, Hultgardhnilsson A (1994), `Fibronectin and the basement membrane components laminin and collagen type-IV influence the phenotypic properties of sub-cultured rat aortic smooth muscle cells differently', Cell Tissue Res, 276, 263± 271. Tsuruma A, Tanaka M, Yamamoto S, Shimomura M (2008), `Control of neural stem cell differentiation on honeycomb films', Colloids Surf A, 313, 536±540. Turner DC, Flier LA, Carbonetto S (1989), `Identification of a cell-surface protein involved in PC12 cell-substratum adhesion and neurite outgrowth on laminin and collagen', J Neurosci, 9, 3287±3296. Valencia S, Gretzer C, Cooper LF (2009), `Surface nanofeature effects on titaniumadherent human mesenchymal stem cells', Int J Oral Maxillofac Implants, 24, 38± 46. Wells N, Baxter MA, Turnbull JE, Murray P, Edgar D, Parry KL, Steele DA, Short RD (2009), `The geometric control of E14 and R1 mouse embryonic stem cell pluripotency by plasma polymer surface chemical gradients', Biomaterials, 30, 1066±1070. Xu CH, Inokuma MS, Denham J, Golds K, Kundu P, Gold JD, Carpenter MK. `Feederfree growth of undifferentiated human embryonic stem cells' (2001), Nat Biotechnol, 19, 971±974. Yang F, Williams CG, Wang DA, Lee H, Manson PN, Elisseeff J (2005), `The effect of incorporating RGD adhesive peptide in polyethylene glycol diacrylate hydrogel on osteogenesis of bone marrow stromal cells', Biomaterials, 26, 5991±5998. Yang XB, Roach HI, Clarke NMP, Howdle SM, Quirk R, Shakesheff KM, Oreffo ROC (2001), `Human osteoprogenitor growth and differentiation on synthetic biodegradable structures after surface modification', Bone, 29, 523±531. Yim EKF, Pang SW, Leong KW (2007), `Synthetic nanostructures inducing differentiation of human mesenchymal stem cells into neuronal lineage', Exp Cell Res, 313, 1820±1829. Yokota T, Oritani K, Mitsui H, Aoyama K, Ishikawa J, Sugahara H, Matsumura I, Tsai S, Tomiyama Y, Kanakura Y, Matsuzawa Y (1998), `Growth-supporting activities of fibronectin on hematopoietic stem progenitor cells in vitro and in vivo: Structural requirement for fibronectin activities of CS1 and cell-binding domains', Blood, 91, 3263±3272.

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Modifying biomaterial surfaces to optimise interactions with bone R . L . S A M M O N S , University of Birmingham, UK

Abstract: Repair and replacement of bone and parts of the skeleton by biomaterials presents many challenges because of the complexity of bone and its functions. In an aging population ever-increasing numbers of prosthetic devices and biomaterials are required. This chapter discusses how surface treatments are currently used to optimise the performance of joint replacements, bone graft substitute materials and in dental implants, and future developments which aim to improve their performance even in patients whose healing capacity is compromised by age-related disorders. Key words: hip replacements, bone graft substitute materials, dental implants, wear, surface roughness, osseointegration.

15.1

Introduction

Repair and replacement of bone and parts of the skeleton by biomaterials present many challenges because bone is a complex tissue with many functions. It is a natural composite in which fracture toughness is conferred by an organic phase consisting mainly of collagen type 1, and hardness by embedded crystals of carbonate-substituted hydroxyapatite. The lamellar structure of mature bone also contributes to its mechanical properties. In Haversian systems mineralised collagen fibres are arranged in concentric cylinders or lamellae surrounding a central capillary. In adjacent lamellae the mineralised collagen fibres are orientated in different directions, a property which enables the bone to resist stresses from different directions. Dense cortical bone provides the strength required to support the body and resist muscular forces, whilst porous trabecular or cancellous bone is responsible for transferring stress to the cortical bone (Fig. 15.1). Medical devices that adhere to bone should disrupt the normal stress strain relation relationships as little as possible to avoid stress shielding. Bone is constantly remodelled throughout the lifetime of an animal in response to physiological changes and fatigue and the partnership between osteoblasts that make bone and osteoclasts that destroy it is finely tuned so that under normal circumstances there is no net loss of bone (Martini, 1988). Biomaterials may impact on these processes via their influence on osteoblasts and osteoclasts. On

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15.1 Photograph of section through a human femur showing porous trabecular (cancellous) and dense cortical bone. Trabeculae cross each other at right angles, following the lines of stress and transfer force from the head of the femur to the cortical bone in the shaft. Inset: (a) arrangement of Haversian systems (osteons) in mature trabecular as well as cortical bone; (b) arrangement of collagen fibres in adjacent osteon lamellae; (c) arrangement of crystals of bone mineral in the gaps between adjacent tropocollagen molecules in collagen fibrils; bundles of fibrils make up collagen fibres. Photograph by the author; inset modified from Cowin et al., 1987.

the other hand, osteoclast cells may also attack the biomaterials surfaces leading to the release of ions that can trigger hypersensitivity reactions (Cadosch et al., 2009a,b). As our understanding of bone cell and molecular biology has increased over the last 20 years our knowledge and understanding of the interactions that occur at the interface between bone and biomaterials has advanced in parallel. Rapid advances have also been made in our understanding of the factors that can limit the life of prosthetic devices in the body and the means to combat these. The demand for devices such as replacement joints increases each year as people live

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15.2 Why we need joint replacements. Photograph of human acetabular cup showing signs of wear on the cartilage (arrows). The inset shows wear on the head of the humerus. Images obtained from anatomical skeleton by the author.

and want to remain active longer yet still continue to suffer from common bone disorders such as osteoarthritis (Fig. 15.2). In the dental field more and more people are demanding dental implants for cosmetic reasons and research is directed towards decreasing treatment time with immediate placement and early loaded implants. We are now seeing the development and application of materials that will actively promote bone healing at the cellular and molecular level (Navarro et al., 2008). Modification of the surface properties of implanted medical devices and the biomaterials from which they are composed is the key to improvements in materials performance which will make them available to immunosuppressed patients or those whose healing capacity is compromised by age-related disorders such as osteoporosis and type II diabetes. Surface modifications are needed: 1. To improve the functional lifetime of an implant by decreasing the rate of degradation due to corrosion and wear. 2. To promote and improve fixation by improving the strength of adhesion to host bone and to accelerate the rate of bonding. 3. To promote osseointegration by influencing the attachment and differentiation of osteoblasts. 4. In the case of resorbable graft materials, to permit the attachment of osteoclasts, formation of resorption lacunae and eventual remodelling. 5. To deliver cells, biomolecules or drugs to the implant site to promote healing and prevent infection.

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15.3 Applications of surface modifications of bone-contacting materials. Redrawn and modified from Lappalainen et al., 2005.

6. To make the device visible using various imaging techniques in order to reveal changes or movement. These are summarised in Fig. 15.3. In this chapter I have attempted to describe the problems and challenges that bone-contacting materials currently face and to discuss how surface treatments are used to improve performance. I then briefly discuss some of the current research that will lead to improvements in performance of bone-contacting biomaterials in this very broad and rapidly developing field. I have divided the devices and biomaterials into three categories: 1. Joint replacements, bone plates, screws and other fixation devices for bone repair and reconstruction (focussing mainly on hip replacements). 2. Bone graft substitute materials for filling cavities in bone and for bone augmentation (focussing mainly on calcium phosphate ceramics). 3. Percutaneous devices that anchor to bone and support an external prosthetic device (focussing on dental implants).

15.2

Joint replacements: successes and challenges

Joint replacement has become a routine operation that improves the quality of life of thousands of people every year. Approximately 250 000 hip replacements are carried out each year in Europe and the USA (Crowninshield et al., 2006). Knee joint replacements come a close second and there are increasing numbers of operations to replace the elbow, shoulder, wrist and hand joints. In 2009, 144,093 hip and knee procedures were recorded in England and Wales alone, according to the National Joint Registry.

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Materials for joint prostheses have been selected because of their relatively high resistance to corrosion and fatigue and because they are comparatively well tolerated by the human body. However, from a material's perspective the body represents a hostile environment where it is subjected to both physical and biological forces (Lappalainen et al., 2005; Yan et al., 2006; Cadosch et al., 2009a,b; which can lead to corrosion and the production of wear particles. Aseptic loosening resulting from osteolysis due to the inflammatory reaction to wear particles and/or ions released into the tissues is a major cause of failure of hip joints. Other causes of failure include stress shielding, dislocation and infection. Aseptic loosening due to wear has been extensively researched (reviewed by Harris, 1995; Agarwal, 2004; Drees et al., 2007; Cadosch et al., 2009b). Particles of polymers and metals released from the articulating surfaces gradually build up in the tissues adjacent to the implant where phagocytosis by macrophages and other cells results in the release of cytokines which stimulate the proliferation and differentiation of osteoclast precursors, their fusion to form osteoclasts leading to resorption of bone. Osteolysis is correlated with the number of particles and is seen when this reaches a critical threshold, estimated as about 1010 particles per gram of tissue (Kobayashi et al., 1997; Kadoya et al., 1998). Particle number appears to be the most important factor although size, shape and composition all influence the extent and characteristics of the inflammatory response (Doorn et al., 1996; Green et al., 1998). A major focus of research has been surface treatments to reduce the production of wear particles.

15.2.1 Methods to reduce wear of articulating joint surfaces In a total hip replacement a metal femoral component articulates with an artificial acetabular cup which, in many designs, is lined with ultrahigh molecular weight polyethylene (UHMWPE), although metal-on-metal and ceramic-on-ceramic joints are also used. Wear particles from UHMWPE used to be a major cause of osteolysis in both hip and knee replacements but the introduction of highly cross-linked UHMWPE in 1998 greatly reduced the wear rate (Galvin et al., 2006; Jacobs et al., 2008). Cross linking is generally achieved by exposure to ionising radiation followed by heating or annealing or Vitamin E stabilisation. Clinical studies report a 50±70% reduction in wear rate compared with earlier materials, although there is quite a lot of variation between different products (Jakofsky, 2008). Metal-on-metal and ceramic-on-ceramic bearing surfaces show a greatly reduced rate of wear and metal-on-metal bearings are currently the only option for surface replacement arthroplasty, where only the head of the femur is replaced, (Jacobs et al., 2009). Wear of a polyethylene liner is caused by abrasion by particles originating from the counter-surface. To reduce the number of particles, the abrasion resistance of metal bearing surfaces can be increased by surface treatments such as C ion implantation (Koseki et al., 2008) and O and N ion implantation in

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titanium and titanium alloys, although this is still experimental (Dearnaley et al., 2007). Another way of increasing hardness is by coating with a layer of biocompatible, wear-resistant material with a very low coefficient of friction. Plasma coating technologies have also been used effectively for this application (Liang, 2004). Coatings include alumina, diamond-like carbon and zirconia and of these, alumina has had the most success. Other materials, particularly diamond-like carbon are not so popular due to problems with delamination or cracking of the coatings in the aqueous body environment (reviewed by Roy and Lee, 2007). `Oxinium', an oxidised form of zirconia with superior hardness properties is used for the condylar parts of hip and knee replacements and can also be used for bone plates, screws and other bone-fixation devices (Good et al., 2005; Hunter et al., 2009; Fig. 15.4). Titanium alloys are now preferred to cobalt chrome molybdenum alloys for many designs of hip and other joint replacements because of their lightness, lower elastic modulus, superior corrosion resistance and better biocompatibility

15.4 Modular designs of cementless hip replacements in which the roughened portion for bone in-growth is confined to the proximal region and in this design achieved by in-cast beads (see also Fig. 15.5). The image on the left shows a SYNERGY TM hip with OXINIUMTM head. Images courtesy of Smith and Nephew, UK.

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(Long et al., 1998; Liu et al., 2004). To improve their tribological performance the bearing surfaces of titanium joints and the entire surface of artificial bones such as the collar bone, may be coated with more wear-resistant materials such as titanium nitride, TiAlN and Al2O3 (Buechal et al., 2004; Yildiz et al., 2009) by plasma nitriding, thin film deposition, and plasma spraying. Physical methods are also used for surface hardening of metals. For example, shot peening, in which compressive strength is imparted to the surface by bombarding it with small spheres, is used for hardening the tapered (ball-attachment) joint of modular hip replacements and to improve the fatigue resistance and fretting wear of hip stems and the fatigue resistance of orthopaedic screws (Progressive SurfaceTM http://www.progressivesurface.com).

15.2.2 Cemented joints: surface treatments to improve cement bonding The surfaces of cemented joint prostheses which come into contact with cement should not strictly be included in this chapter because they are not interacting with bone but are nevertheless of interest because the optimal surface for cemented femoral components of hip replacements is still debated: Lachiewicz et al. (2008) compared cemented hips of the same geometrical design with either smooth polished surfaces (Ra, 0.18±0.3 m) with ones with a roughened surface (Ra, 1.8±2.3 m) for an average of 5 years and showed no difference in performance. They concluded that cementing technique may actually be more important than the surface finish because failure to produce a tight fit to seal in wear particles could result in osteolysis. As with any device, geometry and surface properties must evolve in concert for optimal performance. In collarless, double tapered designs such as the Exeter hip (Stryker, Newbury, UK), the stem is highly polished to allow the hip to subside 1±2 mm in the socket within the cement mantle as it is loaded, increasing the tightness of fit and distributing the load to proximal bone to avoid stress shielding. In an earlier design with the same geometry but a matte finish the abrasive action of the surface against the cement led to the production of enormous numbers of cement particles and resulted in unacceptable numbers of failures due to osteolysis (Shah and Porter, 2005).

15.2.3 Non-cemented joints: wear and surface treatments to improve bone fixation; hydroxyapatite coatings Problems with cement degradation leading to the production of cement wear particles in the 1980s led to the development and use of non-cemented hip joints. Although cements have improved in recent years and cemented hips can be used in patients of all ages, many surgeons prefer to use a press-fit design with a stem which has been roughened to promote bone growth into the pits and crevices of

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the surface and achieve primary fixation by mechanical interlocking. For primary replacements the roughened area is usually confined to the proximal region of the stem of the prosthesis to ensure optimal stress transfer to the host cortical bone and avoid stress shielding (Fig. 15.4). Roughening is achieved by a number of different methods including grit-blasting with aluminium oxide or other abrasives, plasma-spraying a coating of titanium or hydroxyapatite or by casting or sintering porous titanium onto the surface, which in some designs are then additionally coated with hydroxyapatite. Hydroxyapatite (HA) coatings have been used for fixation of titanium cementless hip replacements for over 20 years (reviewed by Naryanan, 2008). Early experiments in animals provided evidence that HA could promote bone formation and bone bonding even in the presence of micromotion or a 0.5 mm gap between bone and the implant (Sùballe et al., 1993). HA also allows remodelling in response to stress around the implant (Chandran et al., 2009). Some coatings consist of a mixture of HA and -tricalcium phosphate: HA is insoluble at physiological pH and provides stability, whilst the more soluble TCP is intended to provide a localised source of calcium and phosphate ions which reprecipitate as carbonate apatite, accelerating mineralisation at the interface. Cells react differently to HA and -TCP: Dos Santos et al. (2009) reported that cells growing on -TCP were most influenced by surface chemistry which reduced adhesion but enhanced differentiation, whereas on HA, surface topography increased differentiation but lowered proliferation. As with porous metals the HA coating is usually confined to the proximal region for optimal stress transfer to remaining trabecular and cortical bone (Huiskes et al., 1993). However, Chandran et al. (2009) reported that Furlong stems, which are coated along their entire length with a 200 m thick, plasmasprayed layer of HA, showed a similar survival rate at 14 years to Charnley cemented prostheses; the overall survival was 93.6% in the Furlong group and 94.8% in the Charnley group. Whether or not the Furlong stems lead to osteolysis in the longer term remains to be seen. HA is an osteoconductive material which is said to be bioactive, forming a direct bond with bone and there is some evidence for epitaxis of bone mineral and biomaterial HA crystals (Bonfield and Luklinska, 1991) although this is difficult to prove in vivo. It is also difficult to prove that the advantages conferred by HA coatings in vivo are due to the chemistry or to the topography of the surface, or even just to a tighter initial press-fit (Wennerberg and Albrektsson, 2009), since the most common method of applying the coating, plasma-spraying, creates a rough osteoconductive porous microstructure which aids fixation by bone ingress and mechanical interlocking in the same way as porous metal does. Porous coatings allow fixation without cement or screws, thus avoiding the problem that screw holes in acetabular cups may provide a route for wear particles to migrate into the surrounding tissues; a study by Coathup et al. (2005) in sheep suggested that HA coatings on acetabular cups

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15.5 Metal-on-metal hip-resurfacing device which replaces the acetabular cup and the head of the femur only. The acetabular cup is bead-coated and a cross section (insert) shows how the beads are cast into the surface, reducing the risk of detachment. Images courtesy of Smith and Nephew.

may be beneficial in keeping wear particles confined to the joint space by producing a good seal to host bone without fibrous encapsulation. But both HAcoated and uncoated porous surfaces were superior to HA-coated or uncoated grit-blasted ones, highlighting the fact that the topography or geometry of the underlying surface is also important. A similar conclusion was reached by Hayakawa et al. (2002), following a comparison between Ca-P coated and uncoated Ti powder blasted implants and Ca-P coated and uncoated Ti beadcoated implants in rabbits; the bead-coated surfaces were superior whether or not they were HA coated although after 12 weeks the highest amount of bone contact occurred with the Ca-P coated beads. Therefore to take advantage of mechanical interlocking as well as HA bonding, some implants are coated with HA-covered microspheres, which in some designs are cast into the surface to reduce the risk of detachment (Fig. 15.5).

15.2.4 Future developments in surface modifications to hip replacements Improving wear resistance Research is continuing to improve the lifetime of joint replacements by reducing wear. In addition to ion implantation, mentioned above, one possibility is to coat the surface with a material with a very low coefficient of friction. These `superlubricious coatings' have a coefficient of friction which is similar to that

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of natural cartilage and can be grafted onto the surface of polyethylene or metal (e.g. cobalt-chrome molybdenum). A promising example is the phospholipid polymer poly(2-methacryloyloxyethyl phosphorylcholine (MPC; Moro et al., 2004; Kyomoto et al., 2009). In addition to its superior mechanical properties, when grafted onto polyethylene particles the MPC did not stimulate the release of cytokines by macrophages and osteoclastogenesis that could lead to osteolysis (Moro et al., 2004). Diamond-like coatings (DLC) have not been extensively exploited for biomedical applications due to problems with the development of internal stress leading to delamination of thick coatings and cracking of thin ones (Erdemir et al., 2006). However, Alakoski et al. (2008) suggest that with the right choice of substrate, thick amorphous DLC could be used for load-bearing applications including the condylar surfaces of hip and knee joints. The most promising test results have been achieved with DLC-coated polyethylene sliding on DLCcoated CoCrMb alloy (Sheeja et al., 2005, cited by Alakoski et al., 2008). Amaral et al. (2008) have shown that nanocrystalline diamond coatings, which combine very low surface roughness, low coefficient of friction and high wear and corrosion resistance, grown by hot-filament chemical vapour deposition are also highly biocompatible, permitting attachment and proliferation of fibroblasts and promoting differentiation of human osteoblast cells. Yang et al. (2009a,b) showed how DLC on surfaces with different properties (grain size, surface roughness and wettability) can determine osteoblast adhesion and proliferation and suggested that these parameters could be controlled to either promote or inhibit bone growth on different aspects of an implant. Liu et al. (2008) reported that bacterial adhesion was inhibited on DLC containing N or Si on stainless steel; such coatings could also be effective in reducing the incidence of infection associated with fracture fixation pins. Modifications to improve fixation and resist infection Plasma-spraying typically produces an HA coating 50±100 microns thick and delamination of such thick coatings has caused concern and limited the use of, for example, plasma-sprayed HA-coated dental implants (discussed below). Alternative coating methods which can produce thinner coatings, including sputter coating (Yang et al., 2005), beam assisted deposition (Bai et al., 2009), electrophoretic deposition, sol-gel and direct laser melting may be used in the future (de Jonge et al., 2008; Paital and Dahotre, 2009; see also Chapter 6). Kumar et al. (2007) reported the use of plasma treatment to hydroxylate the surface of stainless steel rendering it capable of initiating the formation of a thin layer of amorphous HA on the surface when placed in simulated body fluid (an indication of potential bioactivity). The method has the advantage that it can be used on surfaces with complex geometry. However, more work is needed to determine the optimum coating thickness and other properties required to

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enhance osseointegration in different situations (Coelho et al., 2009a; de Jonge et al., 2009). One of the advantages of HA is the ease with which it can be chemically modified by the incorporation of ions which could promote healing or reduce the risk of infection. For example, strontium ions have been shown to inhibit osteoclastogenesis whilst promoting osteoblast differentiation, and use of strontium in the form of strontium ranelate is already a recognised treatment for osteoporosis (Landi et al., 2007, 2008). The release of strontium ions from strontium-substituted HA could therefore be beneficial for fixation in osteoporotic patients; of course this also applies to HA used as a bone graft substitute material, as discussed in the next section and strontium ions have already been incorporated into HA in bone cement for treatment of kyphoplasty (Cheung, 2005). Infection is a rare but very serious problem in joint arthroplasty, resulting in failure of approximately 1% of hip and knee replacements ± a small but very significant proportion considering the enormous number of procedures. Chen et al. (2009), Ruan et al. (2009) and Ando et al. (2010) have developed HA coatings incorporating silver ions which showed antimicrobial activity whilst retaining cytocompatibility. DõÂaz et al. (2009) reported the manufacture of a bactericidal silver-hydroxyapatite nanocomposite. Such coatings could potentially reduce the risk of `early' infections of joint prostheses due to microorganisms contaminating the wound during surgery. Staphylococci are frequently associated with bone infections and infections associated with stainless steel fracture fixation plates. Bare metal surfaces may also be treated to reduce infection: Zhao et al. (2008) reported that implantation of silicon fluoride into stainless steel reduced bacterial surface contamination and implanted surfaces showed resistance to Staphylococcus aureus and S. epidermidis.

15.3

Bone graft substitute materials: successes and challenges

The number of bone graft procedures carried out annually is huge. Giannoudis et al. (2005) reported that approximately 2.2 million operations were carried out world-wide each year. Bone grafting, using a piece of bone from another part of the body (autograft), a donor of the same species (allograft) or another species (xenograft), has a long history (reviewed by Urist et al., 1994). According to de Boer (1988) the first graft procedure was described by the Dutch surgeon, Job Van Meekeren and involved the allegedly successful repair of a traumatic bone defect in a soldier's skull using bone from the cranium of a dog (de Boer reports that the soldier was excommunicated from the church for having an unchristian treatment and he asked to have it removed later but by then the graft had healed!).

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The `gold standard' material for bone replacement today is autologous bone from the patient which, depending on the application, may be obtained from the iliac crest, ribs, chin or ramus of the mandible. Such grafts provide all the factors needed for successful osseointegration: living cells (at least on the surface), growth factors and a bone scaffold. However, harvesting bone for an autograft may be associated with infection, pain and morbidity of the graft site and scarring (Nkenke et al., 2001, 2002, 2004; Giannoudis et al., 2005) and natural bone may not be present in sufficient quantity in children or may be of poor quality in older patients. Donated or allograft bone brings with it the risk of infection, inconsistent quality and performance. On the other hand a bonesubstitute graft material is provided sterile and is potentially available in unlimited quantities, with theoretically consistent properties and is more convenient for the surgeon and less traumatic for the patient. However, all current bone graft substitute materials have some disadvantages. These include unpredictable rates of resorption, difficulty of handling and the inability to be used in load-bearing locations because of their poor mechanical properties. Research is focussing on overcoming these problems as well as augmenting their capacity to stimulate bone formation. If a surgeon chooses not to use autograft or allograft bone, a large variety of alternatives are available (reviewed by Burg, 2000; Moore et al., 2001; Vaccaro et al., 2002; Giannoudis et al., 2005; Oh et al., 2006; Hallman et al., 2009). For different applications they must choose between resorbable materials that are gradually replaced by bone and non-resorbable materials which maintain volume but take longer to heal. For example, for bone augmentation prior to placement of a dental implant a non-resorbable material would be chosen for the `aesthetic zone' at the front of the mouth whilst a resorbable material might be chosen for augmentation in the maxillary sinus where any loss of volume would be invisible (Hallman et al., 2009). Graft materials can be divided into three large groups: 1. Grafts made from animal bone: both deproteinised bone mineral and demineralised bone (collagen-based materials) are used. One of the most extensively researched materials in the first category is Bio-OssÕ (Geistlich, Baden Baden, Germany; Fig. 15.6), which is derived from deproteinised bovine bone and used extensively in dental applications (Stelzle, 2009). 2. Grafts derived from natural sources that resemble bone: materials in this category include BioCoralÕ (Biocoral Inc., Delaware, USA) consisting of a calcium carbonate scaffold, ProOsteonÕ (Interpore Cross International, Irvine, California, USA; Fig. 15.6) in which the coral calcium carbonate skeleton is hydrothermally converted to calcium phosphate (HA and small amounts of -TCP), retaining the bone-like structure of the original material. FRIOS Algipore (Dentsply Friadent, Mannheim, Germany) is similarly produced from the porous calcium carbonate skeleton of a marine alga.

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15.6 Examples of graft materials that mimic bone architecture: (A and B), BioOssÕ (deproteinised bovine bone), Geistlich, Baden Baden, Germany; (C and D), BoneCeramic, 60% HA; 40% -TCP, StraumannÕ, Basel, Switzerland; (E and F), Ossbone Opencell hydroxyapatite, Curasan AG, Kleinostheim, Germany, (G and H), Interpore 200, (now marketed as ProOsteonÕ), Biomet, Interpore Cross, California, USA. The surface of BioOssÕ is most similar to natural bone. Although the architectures of the others are quite similar to bone, the surface topographies are very different. Magnifications: A, C, E and G 50, size bar = 500 m (or E, 100 m). B, D, F and H show the surface of each material 1000; size bar = 10 m.

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15.7 (A) Application of a bone graft for alveolar bone augmentation prior to placement of a dental implant when the bone width is inadequate for successful placement; (B) a bone graft along with a membrane to contain the graft at the graft site is placed over the existing bone; (C) adequate bone width is regenerated for placement of an implant. Images courtesy of BioHorizons, Inc.

3. Totally synthetic materials: this large group includes calcium carbonate, calcium sulphate and calcium phosphate cements and ceramics: especially hydroxyapatite (e.g., OsteoGraf (Dentsply Friadent Ceramed, Mannheim Germany)), beta-tricalcium phosphate (e.g., Cerasorb; Curasan, Kleinestheim, Germany) or a combination of the two, (e.g., BoneCeramic; StraumannÕ, Basel, Switzerland; Fig. 15.6); bioactive glasses (e.g., Norian SRS, Synthes, Inc, West Chester, PA, USA). Several polymers are used either alone or in the form of composites with calcium phosphates or bioactive glass and porous metals, e.g. titanium foam or fibre scaffolds. They include non-resorbable materials including PMMA and derivatives, polyurethane, polyesters; and resorbable ones including polycaprolactone and polylactic and poly(lactic-co-glycolic acid) based materials An example is Immix Extenders (Osteobiologics, Inc, San Antonio, TX, USA) which is a particulate poly(lactic-co-glycolic acid) material used with autograft bone as a graft extender (Agrawal et al., 2001; Oh et al., 2006; Laurencin and Khan, 2009). Some examples of commercial graft materials are shown in Fig. 15.6. Applications include spinal fusion, as a bone filler following the removal of cysts and as a bone extender with autograft for fracture fixation. In the oral cavity bone grafts are used for alveolar ridge augmentation to increase width (Fig. 15.7) and for augmentation of the maxillary sinus to increase bone height prior to placement of implants in the posterior maxilla (Fig. 15.8).

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15.8 Maxillary sinus augmentation. In this application a bone graft is used to compensate for insufficient bone for implant placement in the posterior maxilla. Left: treatment planning: the position of implants to be placed is drawn on the X-ray. A graft material is placed in the maxillary sinus stimulating new bone formation in the cavity and generating adequate bone height for subsequent placement of implants (right). The upper white line indicates the new bone height. Images courtesy of University of Birmingham, School of Dentistry.

15.3.1 Graft topography and architecture The optimum pore size for scaffold materials is about 300 microns (Tsuruga et al., 1997; Moore et al., 2001; Oh et al., 2006) although secondary osteons can form in cavities as small as 50 micron (Itala et al., 2001). Gray et al. (1996) showed that bone forms preferentially in 350 micron wide grooves of variable depth. There is increasing evidence that architecture and geometry in association with growth factors is critically important for osteoinduction (ectopic bone formation in non-bony sites such as muscle), (Kuboki et al., 2001). This is of interest because materials that are osteoinductive also tend to form more bone in bony sites (Habibovik et al., 2007). Experiments by Ripamonti and co-workers (reviewed by Ripamonti, 2004), in which they implanted highly crystalline hydroxyapatite blocks with different sized concavities into baboon muscles showed that bone was induced only inside the cavities and was associated with the concentration of BMP from the tissues on the inner walls of the concavities. They proposed that materials with the right architecture and surface chemistry could be used to induce bone without the need to add exogenous growth factors (as is current practice in some applications (Giannoudis et al., 2005)). However, the mechanisms leading to osteoinduction are still not well understood

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(Habibovik et al., 2007) and further research is needed to understand the physical and biological factors that govern this process.

15.3.2 Organic surface modifications All bone graft materials must be osteoconductive and have interconnecting pores to allow ingress of blood vessels and surrounding bone. Only autograft bone is osteoinductive, although synthetic grafts can be rendered so (as in the case of HA as described above) in combination with purified growth factors or bone marrow aspirate which contains these (Hallman et al., 2009). One of the major uses of bone graft materials is in the repair of non-healing fractures, and to assist the healing process, growth factors or attachment molecules, which promote the differentiation of osteoblasts or recruit osteoblast precursors, may be included with the material or grafted onto the surface. Examples include recombinant human bone morphogenetic proteins (rhBMP-2 and rhBMP-4), osteoprogenitor protein 1 (rhBMP-7) and peptide P15. Further information on BMP and other biological molecules for bone and cartilage engineering can also be found in the review by Ohba et al. (2009). BMPs are used either alone or in combination with autograft, allograft or bone graft substitute materials to promote healing of tibial and scaphoid nonunion fractures and in spinal fusion (reviewed by Garrison et al., 2007). Currently only BMP-2 and BMP-7 are used clinically: BMP-2 promotes the conversion of mesenchymal stem cells to osteoprogenitor cells and is then involved with their differentiation to mature osteoblasts; BMP-7 is involved in differentiation after the conversion stage. In the UK BMP-2, is supplied in a collagen sponge carrier (InductOsTM) by Stryker UK Ltd; BMP-7 in a bovine collagen carrier in granular form (OsigraftÕ), by Wyeth Research (UK) Ltd. Experimental carriers include collagen and degradable polymer (PLGA and PLLA) sponges, microspheres, nanoparticles and nanofibres. In their 2009 review Ziyad et al. conclude that there is currently no single clinically effective delivery system for BMPs and other growth factors. The reason for the caution in approving BMP-treated materials is that the release kinetics of the protein are difficult to control and could potentially provoke ectopic bone formation. In fact there is some evidence from animal experiments that grafts containing rhBMP-2 and titanium coated with rhBMP-2 may perform less well in terms of bone-toimplant contact than controls without it, at least in the later stages of osseointegration (reviewed by Junker et al., 2009). Further work is necessary to establish optimal delivery conditions. The 15 amino acids residues numbering 766±780 of the a1 chain of the type 1 collagen molecule (P15) are responsible for cell attachment and for initiating the sequence of events which ultimately results in bone formation (Thorwarth et al., 2005). In 1996 Oian showed that this 15 amino acid peptide promotes human dermal fibroblast attachment to anorganic bone mineral in a dose dependent

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manner. Recent studies, however, suggest that the peptide may act by promoting osteoblast differentiation rather than cell attachment (Hennessey et al., 2009). Synthetic P15 peptide is marketed as PepGen P-15TM adsorbed onto deproteinised bovine bone OsteografÕ/N or Algipore (DENTSPLY Friadent, Mannheim, Germany) to promote healing of periodontal defects (Yukna et al., 2000) and showed good bone results in maxillary sinus augmentation for dental implant placement (Degidi et al., 2005). The graft material is mixed with blood and packed into the wound which is then covered by a collagen or thin perforated titanium membrane held on with tiny titanium tacks to prevent connective tissue invasion and allow bone healing.

15.3.3 Future developments in bone graft materials Controlling the rate of resorption One of the problems with resorbable graft materials is that they can resorb too quickly and provide no support for bone development. Calcium sulphate hemihydrate (Plaster of Paris) has a long history as a bone graft material. In the body it dissolves completely into calcium and sulphate ions and as it does so it is replaced by a lattice of calcium phosphate which encourages bone formation and angiogenesis (Ricci et al., 2000; Slater et al., 2008). It can be combined with bone and other graft materials and used to deliver antibiotics into the wound site. In 2008 a calcium sulphate/PLLA composite material was developed (Ricci et al., 2008) which resorbs more slowly. This material is already showing some success in in vivo trials where it appears to stimulate more bone formation than the control (Mamidwar et al., 2008a,b; Pecora et al., 2009). Improving the mechanical properties of graft materials Unlike brittle ceramic materials the elastic modulus of some polymers can be much closer to that of bone, and using rapid prototyping techniques they can be manufactured in a variety of different conformations. As composite materials with ceramics, and functionalised with cell attachment peptides and signalling molecules, they show great promise for the future (Stevens, 2008). In the last 10±15 years there has been an explosion of interest in nanophase materials for biological and medical applications. Because of the smaller grain sizes and larger number of grain boundaries nanophase ceramics may also be stronger than conventional materials. Moreover nanophase hydroxyapatite and other ceramics with grain sizes or pore sizes at the nanoscale (< 100 nm) are reported to enhance osteoblast adhesion and function, decrease fibroblast adhesion, and enhance bone remodelling, compared with non-nanophase ceramics (Webster et al., 1999, 2000, 2001; Balasundarum et al., 2006). There are clearly many potential gains to be had from the use of nanophase materials in bone-

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contacting applications but most studies have so far been conducted in vitro with only a few animal studies (Streicher et al., 2007). The future looks very promising for such materials alone or in combination with polymers but much more basic research work needs to be done to ensure that the tiny particles do not cause adverse reactions in the short or long term (Zhang and Webster, 2009). However, `nanorough' and `nanopatterned' surfaces are already being developed for directing osteoblast adhesion (Puckett et al., 2008) and at least one dental implant company is already using such a surface combined with nano-HA, as discussed below. Surface modifications to combat infection Infection is a major problem in bone repair, especially when a trauma wound is heavily contaminated, and in revision surgery especial care must be taken to completely eliminate the infection before replacement of an infected joint prosthesis. Several methods involving bone grafts are used to compact infection, for example antibiotic impregnated cements and ceramic beads are commonly used. As resistance to previously effective antibiotics increases in the general population new agents and methods for combating infection are being sought. Hilpert et al. (2009) have identified cationic peptides with antimicrobial activity which can be tethered to a cellulose support surface without loss of activity. In principle these peptides could be attached to other polymers for a variety of clinical applications.

15.4

Percutaneous bone-anchored devices

This group includes bone-anchored hearing aids, external fracture fixation devices, and dental implants. All these devices transverse the skin and underlying soft tissue and penetrate bone. I have focussed mainly on dental implants because the factors affecting their osseointegration have been studied most extensively.

15.4.1 Dental root implants: successes and challenges Dental implants are used to replace teeth that are absent because of congenital defects or have been lost due to disease or trauma. They present a number of different challenges: not only do they have to successfully osseointegrate with the alveolar bone in order to support the prosthetic tooth, they must also connect the external to the internal environment, as a normal tooth does, and resist infection. The design and surface properties of dental implants continue to evolve in response to demands for implants that osseointegrate more quickly and can be loaded earlier. Another goal is to provide implants which will function even in patients with poor alveolar bone density or whose healing ability is

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compromised. Driven by commercial pressures, dental implant companies are constantly developing new surfaces to try to gain a competitive edge (Table 15.1). However, the competition has also resulted in intensive research which has also contributed to our understanding of the interactions between bone cells and dental implant surfaces. Table 15.1 Evolution of dental implant surfaces from five of the major companies showing the launch of new surfaces in the last 6±7 years Company

Name of implant surface and date of release

Surface treatment Grit Acidblasted etched

Dentsply DPS (1992) Yes Friadent, Yes Mannheim, PlusÕ (2003) Germany http://www.dentsply-friadent.com Astra Tech, MÎlndal, Sweden

TiOBlast (1990)

Yes (TiO2)

Yes OsseoSpeedTM (2004) (TiO2) http://www.astratechdental.com/

Other modification/effect

Yes Yes

As DPS but with pits 3±5 mm diameter; 2±3 mm depth

No Yes (HF)

Incorporation of F ions into Ti oxidea

Nobel Biocare, TiUniteÕ (2000) No Gothenburg, Sweden GroovyTM (2005) No

No

Spark anodisation gives porous surface; P enriched

Straumann, Basel, Switzerland

SLA (1994)

Yes

Yes

SLActive (2005)

Yes

Yes

Biomet 3i, Florida, USA

Osseotite (1996)

No

Yes

Nanotite (2007)

No

Yes

As TiUniteÕ but with 100 m grooves in flanks of threads to improve fixation http://www1.nobelbiocare.com/en/implants-andabutments/concepts/groovy.aspx No

As SLA but modified acid etching confers hydrophilicityb www.straumann.com/ci_ch_presskit_slactive_ids07.pdf - Switzerland

Deposition of nanoscale calcium phosphate crystals on nano-pitted surface http://biomet3i.com/countries/en/directory/Pdf/ART988D_NanoTite_Brochure.pdf a

F incorporation results in a slight increase in surface roughness and height of nanoscale peaks. Wettability is preserved in liquid packaging that also prevents surface contamination with hydrocarbons. b

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15.4.2 Overall design of a dental implant and osseointegration Most dental implants are made of commercially pure titanium (cpTi) because it is malleable enough to be milled to provide the fine screw threads and sharp cutting edge of the self-tapping screw design (as can be seen in Fig. 15.7) and because of its high success rate in osseointegration. The screw threads provide primary stability via mechanical interlocking with bone, aided in many cases by a rough microstructure. This is usually obtained by grit blasting with alumina, titanium dioxide or HA to give 10±100 m cavities which increase the surface area in contact with bone. In many designs a secondary microstructure is added by acid etching or anodic oxidation to provide micron-scale (or in some cases nanoscale) pits and pores on the surface. Some examples of different microstructured dental implant surfaces are shown in Fig. 15.9.

15.9 Examples of microstructured commercially pure titanium dental implant surfaces showing the sizes of surface features relative to attached rat calvarial osteoblast cells. (A) Dentsply Friadent (Mannheim, Germany) PlusÕ ; (B) Dentsply Friadent DPS; (C) Nobel Biocare TiUnite (Gothenburg, Sweden); (D) 3i Biomet OsseotiteÕ (Florida, USA). (A) and (B) are grit blasted and acid etched; (C) is an anodised surface; (D) is acid etched. The spider-like cell morphology on (D) is typical for this surface. Magnification 2000; bar ˆ 10 m. Images by the author.

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15.4.3 Sequence of events in osseointegration Schwartz and Boyan (1994) described the sequence of events and approximate timescale of osseointegration of a dental implant: within seconds of implantation proteins from blood adsorb to the implant surface. Platelets attach to these and to the surface of the implant and a blood clot forms. Undifferentiated mesenchymal cells then migrate to the surface of the implant from the surrounding bone, differentiate into osteoblasts and produce a collagen type I matrix. This then mineralises and bone is formed and remodelled in response to the forces placed upon it. The entire process is orchestrated by endocrine, paracrine and autocrine factors and takes about 21 days up to the start of the remodelling phase.

15.4.4 Influence of surface roughness and topography on osseointegration Surface modifications have dominated dental implant design over the last 10 years (Coelho et al., 2009b). This is because it can be demonstrated that surface roughness influences the sequence of events in bone healing described above at several stages: at the molecular level roughness affects wettability and hence the initial interaction of the surface with proteins and other biomolecules (Rupp et al., 2004). This influences platelet attachment and the blood clotting process. Rough surfaces may enhance blood clot retention by physical entrapment of fibrin molecules and localisation of signalling molecules near the surface of an implant (Davis, 1998; Di Iorio et al., 2005; Fig. 15.10). Osteoprogenitor cells migrate towards the implant surface as fibrinolysis occurs, possibly guided by the fibrin matrix.

15.10 Interactions of blood with `rough' and `smooth' microstructured dental implant surfaces. (A) Dentsply Friadent PlusÕ surface; (B) Dentsply Friadent DPS surface (as in Fig. 15.9). (A) Fibrin forms a three-dimensional `web' on the rough surface trapping red blood cells within it. (B) On the smoother DPS surface the fibrin forms a layer tightly bound to the surface with red blood cells on the outside (Di Iorio et al., 2005). Images provided by Tonino Traini, Dental School, University of Chieti-Pescara, Italy.

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Cells are sensitive to micro- and nanoscale surface topography. Numerous in vitro studies have shown that primary osteoblasts and osteoblast-like cells show a more differentiated osteogenic phenotype on rough surfaces with pits and pores about 1±3 microns in diameter and depth, in comparison with smoother surfaces (Boyan et al., 2001, 2002; Albrektsson et al., 2008). Mineralisation of rat calvarial osteoblasts is enhanced on rough surfaces in vitro (Boyan et al., 2002; Zinger et al., 2005). Wieland et al. (2005) showed that in vitro osteoblast differentiation was enhanced to the greatest extent on surfaces which were both grit-blasted and acid etched, indicating that cells were also influenced by larger sized cavities. In vivo comparisons are rarer but a study in mini-pigs showed that osteogenic factors TGF- 1 and BMP-4 increased at earlier time points adjacent to porous (Nobel Biocare TiUnite) implants, in comparison with smooth machined (Nobel Biocare MKIII) ones (Schierano et al., 2005). Like fibrin, collagen fibres also entangle themselves in rough surfaces in vitro creating a three-dimensional matrix (Fig. 15.11), which could be conducive to bone formation in vivo, but other factors must also be considered. Reports suggest that hydrophilic surfaces may accelerate the early stages of osseointegration (Schwartz et al., 2009). The StraumannÕ (Waldenburg, Switzerland) SLA surface is a rough grit-blasted acid etched surface whose microstructure is quite similar to that of the Friadent plusÕ surface (Fig. 15.9). Recently the company subtly changed the surface chemistry by hydroxylation/hydration making it more hydrophilic and changing the name to SLActiveÕ (Table 15.1). There is in vitro evidence that this surface differentially influences gene expression associated with osteogenesis in vitro in comparison with the original surface

15.11 Cells and extracellular matrix on Dentsply Ankylos implant with Friadent PlusÕ surface. As with fibrin (Fig. 15.10), collagen fibres form a threedimensional matrix on the surface. Rat calvarial osteoblast cells cultured in vitro. Image by the author.

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and it may promote osteogenesis in vivo earlier than its predecessor (Wall et al., 2009) but the long-term success rate in patients is as yet unknown. More recently attention has turned to surfaces with nanoscale features because of their claimed ability to further promote osteogenesis (Sato and Webster, 2004; Zhao et al., 2007), as discussed above. The apparent advantages of nanotopography and nanoscale HA lead to the development and release of 3i Biomet Nanotite implant whose surface has micron-scale pits containing tiny crystals of hydroxyapatite. Whilst there is much evidence to suggest that micro- and nano-rough surfaces promote an osteogenic phenotype in vitro and some evidence that they may accelerate initial bone formation in vivo there may be no advantage in the longterm. At the moment, `there is no evidence showing that any particular type of dental implant has superior long-term success . . . there is a need for more randomised controlled trials with at least 5-year follow up' (Esposito et al., 2005). This does not mean that the surfaces have no influence but simply that there is insufficient evidence for any advantage in the long-term at the moment. There is however some evidence that periodontitis may be associated with rougher implants, possibly due to entrapment of bacteria in surface cavities.

15.4.5 Ion-implantation Two of the goals in dental implantology are to promote faster osseointegration for earlier or immediate loading and to achieve successful retention in the posterior maxilla which tends to have low bone density. The interaction of proteins with surfaces depends on surface chemistry which is determined by the ionic species present on the surface. Cell attachment and subsequent behaviour is mediated via integrin binding to specific protein sequences hence ion implantation can alter cell responses (reviewed by Braceras et al., 2007). For example, Ca2+ enhanced cell adhesion in vitro and osteoconduction as measured by bone-to-implant contact, in rabbits (de Maeztu et al., 2008). In the same article the authors reported that osteoblast cells grown on Ti implanted with CO+ ions produced higher levels of alkaline phosphatase in vitro than controls, suggesting earlier differentiation, and in vivo a higher bone to implant contact was observed in rabbits in areas of poor bone density and in dogs. Similarly, fluoride ions are well known to enhance osteoblast responses to Ti in vitro and to increase bone conduction and adhesion strength in vivo (Ellingsen et al., 1995; Lamolle et al., 2009). This is exploited in Astra-Tech's (MoÈlndal, Sweden) fluoride-impregnated implant surface OsseospeedTM (Table 15.1). However it also has a nanostructured surface which is slightly rougher than untreated samples so it will be difficult to tell whether any benefit is due to topography or chemistry (Wennerberg and Albrektsson, 2009).

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15.4.6 Loss of crestal bone height and `platform switching' It is not unusual to observe a small reduction of bone height of about 1 mm within one year of placement of dental implants (Fig. 15.12). However, it seems that this does not occur when the diameter of the abutment is smaller than that of the implant. This is known as `platform switching' (Lazzara et al., 2006) and is a relatively new concept in implantology but one that has rapidly gained popularity because of its success in preventing crestal bone loss. One explanation is that a biological height of about 2±3 mm of soft tissue is required above the bone or implant surface to form a connective tissue seal against infection and bone resorbs until this height is attained (Kim et al., 2006). Therefore to promote attachment of soft tissue to the implant surface and prevent loss of bone height dental implant companies have adapted their designs. One of the first to do so is BioHorizonsÕ (Birmingham, AL, USA), who have used laser surface engineering to produce 8 and 12 micron microchannels on the collar of their Laser-LokTM implant (Fig. 15.13). It is showing promising results in animal studies and clinical trials with reduced crestal bone loss in comparison with controls (Weiner et al., 2008; Alexander et al., 2009). This is one of the first clinical applications of laser surface engineering and demonstrates the potential of the technique to control tissue growth.

15.4.7 Future developments in the surface treatment of dental implants Biomimetic surfaces Laser surface engineering (as discussed above) offers great potential for the future as it can be used to produce hierarchical multifunctional surfaces capable

15.12 X-ray showing crestal bone loss around an implant after about one year. Image provided by Ben Aghabeigi, Tatum Dental Institute, Birmingham, UK.

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15.13 Collar of a BioHorizons Laser-Lok implant which has been modified by the addition of 8 and 12 micron width microchannels using a laser ablation technique. The 12 micron microchannels are shown magnified 1000 in the insert. This design is intended to prevent crestal bone loss by promoting gingival attachment to the collar. Images courtesy of John L Ricci, PhD (NYU) and BioHorizons, Inc.

of interacting with different biological molecules and tissues (discussed by Kurella and Dahotre, 2005, 2006). The use of laser technology for the formation of multiscale features, microstructures and phase evolution in HA coatings is also an exciting prospect (discussed by Paital and Dahotre, 2008). Wu et al. (2008) describe a simple hydrothermal method for producing hydrophilic nanotitanates on titanium surfaces which favour hydroxyapatite deposition and accelerate cell attachment and proliferation. Others have reported different ways of adapting surfaces to guide cell responses: Vetrone et al. (2008) have made `nanopatterned' titanium surfaces by the use of cocktails of acids and oxidants to create reproducibly sized and shaped nanoscale pits which enable cell adhesion, proliferation and differentiation to be controlled. Rossetti et al. (2005) reported their success in coating Ti surfaces with supported phospholipid bilayers to resemble cell membranes that can also incorporate functional molecules such as cell attachment receptors as a method for increasing the biocompatibility of Ti still further. Maddikeri et al. (2008) are investigating the use of poly(L-lysine)grafted-poly(ethylene glycol) (PLL-g-PEG) polymers which adsorb onto negatively charged biomaterial surfaces and repel cell and protein attachment. RGD sequences can then be grafted onto these to restore cell attachment capability but the surfaces remain resistant to bacterial attachment. As they say in the title of their article, this is `A first step toward cell selective surfaces'.

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Similarly, nanoscale textures may also be useful in the fight against infection: an article by Park et al. (2008) and another recent one by Puckett et al. (2010) suggest that certain nanoscale Ti topographies may reduce bacterial adhesion while promoting bone tissue formation; an exciting prospect. Future developments to improve fixation also include addition of signalling molecules and growth factors such as the TGF- superfamily, rhBMP-2, platelet derived growth factors, insulin-derived growth factor, which all promote osteoblast differentiation, and the addition of cell adhesion molecules for example RGD (arginine, glutamic acid, aspartate), the cell non-specific, integrin-binding adhesion peptide motif, KRSR (lysine, arginine, serine, arginine) (osteoblast-specific, proteoglycan binding) and others (Garcia and Reyes, 2005; Bagno et al., 2007; Schliephake et al., 2008, reviewed by de Jonge, 2009). De Jonge et al. (2008) describe a very interesting application of the electrospray deposition technique to immobilise alkaline phosphatase enzyme onto the surface of titanium. Another area likely to expand in the future is incorporating drugs into implant surfaces. This has already been tried in the case of bisphosphonates to promote increased bone density (Le GueÂhennec et al., 2007; MendoncËa et al., 2008). Modifying the stiffness of implants The elastic modulus of cpTi is not a perfect match for that of bone and that of titanium alloys can be better adapted. Recently implants made by direct laser sintering of titanium alloy particles have been introduced onto the market (for example, TixosÕ, Leader Novaxa, Italy). Unlike a natural tooth root, a dental implant is not lined by a periodontal ligament which transmits and cushions the biting force. The implant may induce stress shielding or place unnatural loading force on the bone. Using direct laser fabrication, by modifying the size and packing of the alloy particles in the melt bed, it is theoretically possible to form implants with graded elasticity for a more natural transfer of stress from implant to bone (Hollander et al., 2006).

15.5

Summary and conclusions

Surface properties determine the success or failure of bone-contacting biomaterials via their influence on wear properties, biomolecular and cell interactions and resistance to infection. They can influence the rate of the healing process and osseointegration and determine the long-term success of the implant. Current research is leading to the development of implants that are truly biomimetic in terms of both their bulk and surface properties that will enable them to be used even in patients whose immune status and healing ability are compromised.

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Sources of further information and advice

In this chapter I have attempted to provide an overview of the subject of surface treatments of biomaterials for bone-contacting applications but it is impossible to provide adequate coverage of all aspects of this very broad and rapidly developing field. The reader is advised to consult the many excellent reviews on specialist topics, some of which I have referred to in the text. In addition, the following web pages may be useful for finding additional information. · The Ortho Supersite Free website providing clinical news articles, comment and peer reviewed articles related to orthopaedics (http://www.orthosupersite.com/) · International Bone and Mineral Society A useful forum for anyone interested in bone and mineral research. In addition the Society provides free access to `BoneKey' a valuable resource of peer reviewed on-line articles on bone-related topics (http://www.bonekeyibms.org/) · The Cochrane Collaboration: Cochrane Reviews A source of systematic reviews of clinical trials; accessible via higher education establishments in the UK (http://www.cochrane.org/reviews/) · The National Joint Registry Collects information on all hip and knee replacement operations and monitors the performance of replacement hip and knee joints (implants). Also provides links to similar data bases worldwide (http://www.njrcentre.org.uk/njrcentre/ default.aspx) · Orthopaedic web links (OWL) Free membership of OWL gives access to directory of orthopaedic resources (http://www.orthopaedicweblinks.com/) · Hip arthroplasty A useful web site providing images of normal and abnormal appearances following hip arthroplasty as well as background information on pathology and imaging techniques (http://www.gentili.net/thr) · US Food and Drug Administration A wealth of information including guidance on medical devices, consumer resources; regulations and approvals, etc. (http://www.fda.gov/default.htm)

15.7

References

Agarwal S (2004), `Osteolysis-basic science, incidence and diagnosis', Current Orthopaedics, 18, 220±231. Agrawal C M, Ray R B (2001), `Biodegradable polymeric scaffolds for musculoskeletal tissue engineering', Journal of Biomedical Materials Research, 55, 141±150. Alakoski E, Veli-Matti T, Soininena A, Konttinen Y T (2008), `Load bearing biomedical applications of diamond-like carbon coatings', The Open Orthopaedics Journal, 2, 43±50.

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

Albrektsson T, Sennerby L, Wennerberg A (2008), `State of the art of oral implants', Periodontology 2000, 47, 15±26. Alexander H, Ricci J L, Hrico G J (2009), `Mechanical basis for bone retention around dental implants', Journal of Biomedical Materials Research Part B-Applied Biomaterials, 88B, 306±311. Amaral M, Gomes P S, Lopes M A, Santos J D, Silva R F, Fernandes M H (2008), `Nanocrystalline diamond as a coating for joint implants: Cytotoxicity and biocompatibility assessment', Journal of Nanomaterials. Article number: 894352. Available at http://www.hindawi.com/journals/jnm/2008/894352.html. Ando Y, Miyamoto H, Noda I, Sakurai N, Akiyama T, Yonekura Y, Shimazaki T, Miyazaki M, Mawatari M, Hotokebuchi T (2010), `Calcium phosphate coating containing silver shows high antibacterial activity and low cytotoxicity and inhibits bacterial adhesion', Materials Science & Engineering C-Materials for Biological Applications, 30, 175±180. Bagno A E A (2007), `Human osteoblast-like cell adhesion on titanium substrates covalently functionalized with synthetic peptides', Bone, 40 693±699. Bai X, Sandukas S, Appleford M R, Ong J L, Rabiei A (2009), `Deposition and investigation of functionally graded calcium phosphate coatings on titanium', Acta Biomaterialia, 5, 3563±3572. Balasundaram G, Webster T J (2006), `A perspective on nanophase materials for orthopaedic applications', Journal of Materials Chemistry, 16, 3737±3745. Bonfield W, Luklinska Z B (1991), `High-resolution electronic microscopy of a bone implant interface'. In: Davis J E, editor, The Bone±Biomaterial Interface, University of Toronto Press. Boyan B D, Lohmann C H, Dean D D, Sylvia V L, Cochran D L, Schwartz Z (2001), `Mechanisms involved in osteoblast response to implant surface morphology', Annual Review of Materials Research, 31, 357±371. Boyan B D, Bonewald L F, Paschalis E P, Lohmann C H, Rosser J, Cochran D L, Dean D D, Schwartz Z, Boskey a L (2002), `Osteoblast-mediated mineral deposition in culture is dependent on surface microtopography', Calcified Tissue International, 71, 519±529. Braceras I, Alava J I, Goikoetxea L, De Maeztu M A, Onate J I (2007), `Interaction of engineered surfaces with the living world: Ion implantation vs. Osseointegration', Surface & Coatings Technology, 201, 8091±8098. Buechel F F, Buechel F F, Helbig T E, D'alessio J, Pappas M J (2004), `Two- to 12-year evaluation of cementless Buechel±Pappas total hip arthroplasty', The Journal of Arthroplasty, 19, 1017±1027. Burg K J L, Porter S, Kellam J F (2000), `Biomaterial developments for bone tissue engineering', Biomaterials, 21, 2347±2359. Cadosch D, Chan E, Gautsch O, Filgueira L (2009a), `Bio-corrosion of stainless steel by osteoclasts ± in vitro evidence', Journal of Orthopaedic Research, 27, 841±846. Cadosch D, Chan E, Gautschi O P, Filgueira L (2009b), `Metal is not inert: Role of metal ions released by biocorrosion in aseptic loosening ±current concepts', Journal of Biomedical Materials Research Part A, 91A, 1252±1262. Chandran P, Azzabi M, Miles J, Andrews M, Bradley J (2009), `Furlong hydroxyapatitecoated hip prosthesis vs. the Charnley cemented hip prosthesis', The Journal of Arthroplasty, 25, 52±57. Chen Y K, Zheng X B, Xie Y T, Ji H, Ding C X (2009), `Antibacterial properties of vacuum plasma sprayed titanium coatings after chemical treatment', Surface & Coatings Technology, 204, 685±690.

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with bone

393

Cheung K M C, Lu W W, Luk K D K, Wong C T, Chan D, Shen J X, Qiu G X, Zheng Z M, Li C H, Liu S L, et al. (2005), `Vertebroplasty by use of a strontium-containing bioactive bone cement', Spine, 30, S84±S91. Coathup M J, Blackburn J, Goodship A E, Cunningham J L, Smith T, Blunn G W (2005), `Role of hydroxyapatite coating in resisting wear particle migration and osteolysis around acetabular components', Biomaterials, 26, 4161±4169. Coelho P G, Cardaropoli G, Suzuki M, Lemons J E (2009a), `Early healing of nanothickness bioceramic coatings on dental implants. An experimental study in dogs', Journal of Biomedical Materials Research Part B-Applied Biomaterials, 88B, 387±393. Coelho P G, Granjeiro J M, Romanos G E, Suzuki M, Silva N R F, Cardaropoli G, Thompson V P, Lemons J E (2009b), `Basic research methods and current trends of dental implant surfaces', Journal of Biomedical Materials Research Part B-Applied Biomaterials, 88B, 579±596. Cowin SC, van Buskirk WC, Ashman RB. (1987) `Properties of bone'. In: Skalak R, Chien S, editors. Handbook of Bioengineering. New York: McGraw-Hill, pp. 2.1± 2.27. Crowninshield R D, Rosenberg A G, Sporer S M (2006), `Changing demographics of patients with total joint replacement', Clinical Orthopaedics and Related Research, 266±272. Davis J E (1998), `Mechanisms of endosseous integration', International Journal of Prosthodontics, 11, 391±401. Degidi M, Scarano A, Lezzi G, Orsini G, Perrotti V, Strocchi R, Piattelli A. (2005), `Maxillary sinus augmentation using a synthetic cell-binding peptide: A histologic and transmission electron microscopy case study in man', Implant Dentistry, 14, 371±377. De Jonge L T, Leeuwenburgh S C G, Wolke J G C, Jansen J A (2008), `Organic-inorganic surface modifications for titanium implant surfaces', Pharmaceutical Research, 25, 2357±2369. De Jonge L T, Leeuwenburgh S C G, Van Den Beucken J, Wolke J G C, Jansen J A (2009), `Electrosprayed enzyme coatings as bioinspired alternatives to bioceramic coatings for orthopedic and oral implants', Advanced Functional Materials, 19, 755±762. De Maeztu M A, Braceras I, Alava J I, Gay-Escoda C (2008), `Improvement of osseointegration of titanium dental implant surfaces modified with Co ions: A comparative histomorphometric study in beagle dogs', International Journal of Oral and Maxillofacial Surgery, 37, 441±447. Dearnaley G, Asher J, Peacock a T, Allen S J, Watkins R E J (2007), `The use of thin layer activation to evaluate ion beam surface treatments of orthopaedic implant materials', Surface & Coatings Technology, 201, 8070±8075. De Boer H H (1988), `The history of bone-grafts', Clinical Orthopaedics and Related Research, 292±298. Di Iorio D, Traini T, Degidi M, Caputi S, Neugebauer J, Piattelli A (2005), `Quantitative evaluation of the fibrin clot extension on different implant surfaces: an in vitro study', Journal of Biomedical Materials Research Part B: Applied Biomaterials, 74, 636±642. DõÂaz M, Barba F, Miranda M, GuitiaÂn F, Torrecillas R, Moya J S (2009), `Synthesis and antimicrobial activity of a silver-hydroxyapatite coated nanocomposite', Journal of Nanomaterials 2009, 1±6. Doorn P F, Campbell P A, Amstutz H C (1996), `Metal versus polyethylene wear

ß Woodhead Publishing Limited, 2011

394

Surface modification of biomaterials

particles in total hip replacements', Clinical Orthopaedics and Related Research, S206±S216. Dos Santos E A, Farina M, Soares G A, Anselme K (2009), `Chemical and topographical influence of hydroxyapatite and beta-tricalcium phosphate surfaces on human osteoblastic cell behavior', Journal of Biomedical Materials Research Part A, 89A, 510±520. Drees P, Eckardt A, Gay R E, Gay S, Huber L C (2007), `Mechanisms of disease: molecular insights into aseptic loosening of orthopedic implants', Nature Clinical Practice Rheumatology, 3, 165±171. Ellingsen J E (1995), `Pre-treatment of titanium implants with fluoride improves their retention in bone', Journal of Materials Science: Materials in Medicine, 6, 749± 753. Erdemir A, Donnet C (2006), `Tribology of diamond-like carbon films: Recent progress and future prospects', Journal of Physics D-Applied Physics, 39, R311±R327. Esposito M, Coulthard P, Thomsen P, Worthington H V (2005), `Interventions for replacing missing teeth: Different types of dental implants', Cochrane Database of Systematic Reviews. Galvin A, Kang L, Tipper J L, Stone M H, Ingham E, Jin Z M, Fisher J (2006), `Wear of crosslinked polyethylene under different tribological conditions', Journal of Materials Science: Materials in Medicine, 17, 243±253. Garcia A J, Reyes C D (2005), `Bio-adhesive surfaces to promote osteoblast differentiation and bone formation', Journal of Dental Research, 84, 407±413. Garrison K R, Donell S, Ryder J, Shemilt I, Mugford M, Harvey I, Song F (2007), `Clinical effectiveness and cost-effectiveness of bone morphogenetic proteins in the non-healing of fractures and spinal fusion: a systematic review', Health Technology Assessment, 11, 1±150. Giannoudis P V, Dinopoulos H, Tsiridis E (2005), `Bone substitutes: an update', InjuryInternational Journal of the Care of the Injured, 36, 20±27. Good V, Widding K, Hunter G, Heuer D (2005), `Oxidized zirconium: a potentially longer lasting hip implant', Materials & Design, 26, 618±622. Gray C, Boyde A, Jones S J (1996), `Topographically induced bone formation in vitro: Implications for bone implants and bone grafts', Bone, 18, 115±123. Green T R, Fisher J, Stone M, Wroblewski B M, Ingham E (1998), `Polyethylene particles of a ``critical size'' are necessary for the induction of cytokines by macrophages in vitro', Biomaterials, 19, 2297±2302. Habibovik P, De Groot K (2007), `Osteoinductive biomaterials ± properties and relevance in bone repair', Journal of Tissue Engineering and Regenerative Medicine, 1, 25±32. Hallman M, Mordenfeld A, Strandkvist T (2009), `Bone replacement following dental trauma prior to implant surgery ± present status', Dental traumatology, 25, 2±11. Harris W (1995), `The problem is osteolysis', Clinical Orthopaedics and Related Research, 311, 46±53. Hayakawa T, Yoshinari M, Kiba H, Yomamoto H, Nemoto K, Jansen J A (2002), `Trabecular bone response to surface roughened and calcium phosphate (Ca-P) coated titanium implants', Biomaterials, 23, 1025±1031. Hennessy K M, Pollot B E, Clem W C, Phipps M C, Sawyer a A, Culpepper B K, Bellis S L (2009), `The effect of collagen I mimetic peptides on mesenchymal stem cell adhesion and differentiation, and on bone formation at hydroxyapatite surfaces', Biomaterials, 30, 1898±1909. Hilpert K, Elliott M, Jenssen H, Kindrachuk J, Fjell C D, Korner J, Winkler D F H, Weaver L L, Henklein P, Ulrich A S, et al. (2009), `Screening and characterization

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with bone

395

of surface-tethered cationic peptides for antimicrobial activity', Chemistry & Biology, 16, 58±69. Hollander D A, Von Waltera M, Wirtzb T, Selleia R, Schmidt-Rohlfing B, Paar O, Erlia H J (2006), `Structural, mechanical and in vitro characterization of individually structured Ti-6Al-4V produced by direct laser forming', Biomaterials, 27, 955±963. Huiskes R, Weinans H (1993), Biomechanical aspects of hydroxyalapatite coatings on femoral hip prostheses. In Hydroxylapatite Coatings in Orthopaedic Surgery. Edited by Geesink R G T, Manley M T. New York, Raven Press, 63±80. Hunter G, Asgian C M, Hines G L (2009), Method of Surface Oxidizing Zirconium and Zirconium Alloys and Resulting Product. Smith and Nephew Inc. Itala A I, Ylanen H O, Ekholm C, Karlsson K H, Aro H T (2001), `Pore diameter of more than 100 m is not a requisite for bone ingrowth in rabbits', Journal of Biomedical Materials Research, 58, 679±683. Jacobs J J, Campbell P A, Konttinen Y T, Implant Wear Symposium B (2008), `How has the biologic reaction to wear particles changed with newer bearing surfaces?' Journal of the American Academy of Orthopaedic Surgeons, 16, S49±S55. Jacobs J J, Urban R M, Hallab N J, Skipor A K, Fischer A, Wimmer M A (2009), `Metalon-metal bearing surfaces', Journal of the American Academy of Orthopaedic Surgeons, 17, 69±76. Jacofsky D J (2008), `Highly cross-linked polyethylene in total knee arthroplasty: in the affirmative', The Journal of Arthroplasty, 23, 28±30. Junker R, Dimakis A, Thoneick M, Jansen J A (2009), `Effects of implant surface coatings and composition on bone integration: a systematic review', Clinical Oral Implants Research, 20 (Suppl. 4), 185±206. Kadoya Y, Kobayashi A, Ohashi H (1998), `Wear and osteolysis in total joint replacements', Acta Orthopaedica Scandinavica, 69, 1±16. Kim H, Murakami H, Chehroudi B, Textor M, Brunette D M (2006), `Effects of surface topography on the connective tissue attachment to subcutaneous implants', International Journal of Oral and Maxillofacial Implants, 21, 354±429. Kobayashi A, Freeman M A R, Bonfield W, Kadoya Y, Yamac T, Alsaffar N, Scott G, Revell P A (1997), `Number of polyethylene particles and osteolysis in total joint replacements ± a quantitative study using a tissue-digestion method', Journal of Bone and Joint Surgery ± British Volume, 79B, 844±848. Koseki H, Shindo H, Furuichi I, Baba K (2008), `Carbon-ion implantation improves the tribological properties of Co-Cr-Mo alloy against ultra-high molecular weight polyethylene', Surface and Interface Analysis, 40, 1278±1283. Kuboki Y, Jin Q, Takita H (2001) `Geometry of carriers controlling phenotypic expression in BMP-induced osteogenesis and chondrogenesis', Journal of Bone and Joint Surgery, 83, 105±115. Kumar S, Simpson D, Smart R S C (2007), `Plasma processing for inducing bioactivity in stainless steel orthopaedic screws', Surface and Coatings Technology, 202, 1242± 1246. Kurella A, Dahotre N B (2005), `Review paper: Surface modification for bioimplants: the role of laser surface engineering', Journal of Biomaterials Applications, 20, 5±50. Kurella A, Dahotre N B (2006), `Laser induced hierarchical calcium phosphate structures', Acta Biomaterialia, 2, 677±683. Kyomoto M, Moro T, Iwasaki Y, Miyaji F, Kawaguchi H, Takatori, Y, Nakamura K, Ishihara K (2009), `Superlubricious surface mimicking articular cartilage by grafting poly(2-methacryloyloxyethyl phosphorylcholine) on orthopaedic metal bearings', Journal of Biomedical Materials Research A, 91, 730±741.

ß Woodhead Publishing Limited, 2011

396

Surface modification of biomaterials

Lachiewicz P F, Kelly S S, Soileau E S (2008), `Survival of polished compared with precoated roughened cemented femoral components. A prospective, randomized study', Journal of Bone and Joint Surgery of America, 90, 1457±1463. Lamolle S F, Monjo M, Rubert M, Haugen H J, Lyngstadaas S P, Ellingsen J E (2009), `The effect of hydrofluoric acid treatment of titanium surface on nanostructural and chemical changes and the growth of MC3T3-E1 cells', Biomaterials, 30, 736±742. Landi E, Tampieri A, Celotti G, Sprio S, Sandri M, Logroscino G (2007), `Sr-substituted hydroxyapatites for osteoporotic bone replacement', Acta Biomaterialia, 3, 961± 969. Landi E, Sprio S, Sandri M, Celotti G, Tampieri A (2008), `Development of Sr and Co3 co-substituted hydroxyapatites for biomedical applications', Acta Biomaterialia, 4, 656±663. Lappalainen R, Santavirta S S (2005), `Potential of coatings in total hip replacement', Clinical Orthopaedics and Related Research, 72±79. Laurencin C T, Khan Y (2009), `Bone graft substitute materials', eMedicine, available at: http://emedicine.medscape.com/article/1230616-overview. Lazzara R J, Porter S S (2006), `Platform switching: a new concept in implant dentistry for controlling postrestorative crestal bone levels', International Journal of Periodontics & Restorative Dentistry, 26, 9±17. Le Guehennec L, Soueidan A, Layrolle P, Amouriq Y (2007), `Surface treatments of titanium dental implants for rapid osseointegration', Dental Materials, 23, 844±854. Liang H, Shi B, Fairchild A, Cale T (2004), `Applications of plasma coatings in artificial joints: an overview', Vacuum, 73, 317±326. Liu C, Zhao Q, Liu Y, Wang S, Abel E W (2008), `Reduction of bacterial adhesion on modified dlc coatings', Colloids and Surfaces B-Biointerfaces, 61, 182±187. Liu X Y, Chu P K, Ding C X (2004), `Surface modification of titanium, titanium alloys, and related materials for biomedical applications', Materials Science & Engineering R-Reports, 47, 49±121. Long M, Rack H J (1998), `Titanium alloys in total joint replacement ± a materials science perspective', Biomaterials, 19, 1621±1639. Maddikeri R R, Tosatti S, Schuler M, Chessari S, Textor M, Richards R G, Harris L G (2008), `Reduced medical infection related bacterial strains adhesion on bioactive rgd modified titanium surfaces: A first step', Journal of Biomedical Materials Research Part A, 84A, 425±435. Mamidwar S, Ricci J, Alexander H (2008a), `Timed-release calcium sulfate ceramic nanocomposites as bone graft substitutes', Bioceramics, Vol 20, Pts 1 and 2, 361± 363, 391±394. Mamidwar S, Weiner M, Alexander H, Ricci J (2008b), `In vivo bone response to calcium sulfate/poly l-lactic acid composite', Implant Dentistry, 17, 208±212. Martini F H (1988), Fundamentals of Anatomy and Physiology, 4th edn, Prentice Hall International, Engelwood Cliffs, NJ. MendoncËa G, MendoncËa D, Araga F J L, Cooper L F (2008), `Advancing dental implant surface technology ± from micron to nanotopography', Biomaterials, 29, 3822± 3835. Moore W R, Graves S E, Bain G I (2001), `Synthetic bone graft substitutes', ANZ Journal of Surgery, 71, 354±361. Moro T, Takatori T, Ishihara K, Konno T, Takigawa Y, Matsushita T, Chung U I, Nakamura K, Kawaguchi H (2004), `Surface grafting of artificial joints with a biocompatible polymer for preventing periprosthetic osteolysis', Nature Materials, 3, 829±836.

ß Woodhead Publishing Limited, 2011

Modifying biomaterial surfaces to optimise interactions with bone

397

Narayanan R, Seshadri S K, Kwon T Y, Kim K H (2008), `Calcium phosphate-based coatings on titanium and its alloys', Journal of Biomedical Materials Research Part B-Applied Biomaterials, 85B, 279±299. Navarro M, Michiardi A, Castano O, Planell J A (2008), `Biomaterials in orthopaedics', Journal of the Royal Society Interface, 5, 1137±1158. Nkenke E, Schultze-Mosgau S, Radespiel-Troger M, Kloss F, Neukam F W (2001), `Morbidity of harvesting of chin grafts: a prospective study', Clinical Oral Implants Research, 12, 495±502. Nkenke E, Radespiel-Troger M, Wiltfang J, Schultze-Mosgau S, Winkler G, Neukam F W (2002), `Morbidity of harvesting of retromolar bone grafts: a prospective study', Clinical Oral Implants Research, 13, 514±521. Nkenke E, Weisbach V, Winckler E, Kessler P, Schultze-Mosgau S, Wiltfang J, Neukam F W (2004), `Morbidity of harvesting of bone grafts from the iliac crest for preprosthetic augmentation procedures: A prospective study. ` Journal of Oral and Maxillofacial Surgery, 33, 157±163. Oh S, Oh N, Appleford M, Ong J L (2006), `Bioceramics for tissue engineering applications ± a review', American Journal of Biochemistry and Biotechnology, 2, 49±56. Ohba S, Yano, F, Chung, U (2009) `Tissue engineering of bone and cartilage', IBMS BoneKEy 6, 405±419. Available at: http://www.bonekey-ibms.org/cgi/content/full/ ibmske;6/11/405 accessed 3/10/2010. Paital S R, Dahotre N B (2008), `Review of laser based biomimetic and bioactive Ca-P coatings', Materials Science and Technology, 24, 1144±1161. Paital S R, Dahotre N B (2009), `Calcium phosphate coatings for bio-implant applications: Materials, performance factors, and methodologies', Materials Science & Engineering R-Reports, 66, 1±70. Park M R, Banks M K, Applegate B, Webster T J (2008), `Influence of nanophase titania topography on bacterial attachment and metabolism', International Journal of Nanomedicine, 3, 497±504. Pecora G E, Ceccarelli R, Bonelli M, Alexander H, Ricci J L (2009), `Clinical evaluation of laser microtexturing for soft tissue and bone attachment to dental implants', Implant Dentistry, 18, 57±61. Puckett S, Pareta R, Webster T J (2008), `Nano rough micron patterned titanium for directing osteoblast morphology and adhesion', International Journal of Nanomedicine, 3, 229±241. Puckett S D, Taylor E, Raimondo T, Webster T J (2010), `The relationship between the nanostructure of titanium surfaces and bacterial attachment', Biomaterials, 31, 706± 713. Ricci J L, Alexander H, Hollander B (2008), Time release calcium sulfate matrix for bone augmentation. Patent No. US 6770695. Available at ISI>:// BIOSIS:PREV200400354890. Ricci J L, Alexander H, Nadkarni P, Hawkins M, Turner J, Rosenblum S, Brezenoff L, De Leonardis D, Pecora G E (2000), Biological mechanisms of calcium sulfate replacement by bone. In Bone engineering. Edited by Davies Je. Toronto, Em Squared Inc., 332±344. Ripamonti U (2004), `Soluble, insoluble and geometric signals sculpt the architecture of mineralized tissues', Journal of Cellular and Molecular Medicine, 8, 169±180. Rossetti F F, Bally M, Michel R, Textor M, Reviakine I (2005), `Interactions between titanium dioxide and phosphatidyl serine-containing liposomes: formation and patterning of supported phospholipid bilayers on the surface of a medically relevant

ß Woodhead Publishing Limited, 2011

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Surface modification of biomaterials

material', Langmuir, 21, 6443±6450. Roy R K, Lee K R (2007), `Biomedical applications of diamond-like carbon coatings: a review', Journal of Biomedical Materials Research Part B-Applied Biomaterials, 83B, 72±84. Ruan H J, Fan C Y, Zheng X B, Zhang Y, Chen Y K (2009), `In vitro antibacterial and osteogenic properties of plasma sprayed silver-containing hydroxyapatite coating', Chinese Science Bulletin, 54, 4438±4445. Rupp F, Scheideler L, Rehbein D, Axmann D, Geis-Gerstorfer J (2004), `Roughness induced dynamic changes of wettability of acid etched titanium implant modifications', Biomaterials, 25, 1429±1438. Sato M, Webster T J (2004), `Nanobiotechnology: Implications for the future of nanotechnology in orthopaedic applications', Medical Devices, 1, 105±114. Schierano G, Canuto R A, Navone R, Peirone B, Martinasso G, Pagano M, Maggiora M, Manzella C, Easton M, Davit A, Trombetta A, Amedeo S, Biolatti B, Carossa S, Preti G (2005), `Biological factors involved in the osseointegration of oral titanium implants with different surfaces: A pilot study in minipigs', Journal of Periodontology, 76, 1710±1720. Schliephake H, Scharnweber D (2008), `Chemical and biological functionalization of titanium for dental implants', Journal of Materials Chemistry, 18, 2404±2414. Schwartz Z, Boyan B D (1994), `Underlying mechanisms at the bone biomaterial interface', Journal of Cellular Biochemistry, 56, 340±347. Schwarz F, Wieland M, Schwartz Z, Zhao G, Rupp F, Geis-Gerstorfer J, Schedle A, Broggini N, Bornstein M M, Buser D, et al. (2009), `Potential of chemically modified hydrophilic surface characteristics to support tissue integration of titanium dental implants', Journal of Biomedical Materials Research Part B-Applied Biomaterials, 88B, 544±557. Shah N, Porter M (2005), `Evolution of cemented stems', available at: http:www.orthosupersite.com Sheeja D, Tay B K, Nung L N (2005), `Tribological characterization of surface modified UHMWPE against DLC-coated Co-Cr-Mo', Surface Coating Technology, 190, 231±237. Slater N, Dasmah A, Sennerby L, Hallman M, Piattelli A, Sammons R (2008), `Backscattered electron imaging and elemental microanalysis of retrieved bone tissue following maxillary sinus floor augmentation with calcium sulphate', Clinical Oral Implants Research, 19, 814±822. Sùballe K, Hansen E S, Rasmussen H B, Bunger C (1993), Fixation of porous vs hacoated implants. In Hydroxylapatite Coatings in Orthopaedic Surgery. Edited by Geesink R G T, Manley M T. New York, Raven Press, 107±136. Stelzle F (2009), `Clinical outcomes of sinus floor augmentation for implant placement using autogenous bone or bone substitutes: A systematic review', Clinical Oral Implants Research, 20(Suppl.4), 124±133. Stevens (2008), `Biomaterials for bone tissue engineering', Materials Today, 11, 18±25. Streicher R M, Schmidt M, Fiorito S (2007), `Nanosurfaces and nanostructures for artificial orthopedic implants', Nanomedicine, 2, 861±874. Thorwarth M, Schultze-Mosgau S, Wehrhana F, Kessler P, Srour S, Wiltfang J, Schlegel K (2005), `Bioactivation of an anorganic bone matrix by p-15 peptide for the promotion of early bone formation', Biomaterials, 26, 5648±5657. Tsuruga E, Takita H, Itoh H, Wakisaka Y, Kuboki Y (1997), `Pore size of porous hydroxyapatite as the cell-substratum controls BMP-induced osteogenesis', Journal of Biochemistry, 121, 317±324.

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399

Urist M R, Burwell B T, Geoffrey R, Bone Grafts, Derivatives & Substitutes. Oxford, Butterworth-Heinemann Ltd, 1994. Vaccaro A R, Chiba K, Heller J G, Patel T C, Thalgott J S, Truumees E, Fischgrund J S, Craig M R, Berta S C, Wang J C (2002), `Bone grafting alternatives in spinal surgery', The Spine Journal, 2, 206±215. Vetrone F, Variola F, De Oliveira P T, Zalzal S F, Yi J H, Sam J, Bombonato-Prado K F, Sarkissian A, Perepichka D F, Wuest J D, et al. (2009), `Nanoscale oxidative patterning of metallic surfaces to modulate cell activity and fate', Nano Letters, 9, 659±665. Wall I, Donos N, Carlqvist K, Jones F, Brett P (2009), `Modified titanium surfaces promote accelerated osteogenic differentiation of mesenchymal stromal cells in vitro', Bone, 45, 17±26. Webster T J, Siegel R W, Bizios R (1999), `Osteoblast adhesion on nanophase ceramics', Biomaterials 20, 1221±1227. Webster T J, Ergun C, Doremus R H, Siegel R W, Bizios R (2000), `Enhanced functions of osteoblasts on nanophase ceramics', Biomaterials, 21, 1803±1810. Webster T J, Siegel R W, Bizios R (2001), `Nanoceramic surface roughness enhances osteoblast and osteoclast functions for improved orthopaedic/dental implant efficacy', Scripta Materia, 44, 1639±1642. Weiner S, Simon J, Ehrenberg D S, Zweig B, Ricci J L (2008), `The effects of laser microtextured collars upon crestal bone levels of dental implants', Implant Dentistry, 17, 217±224. Wennerberg A, Albrektsson T (2009), `Structural influence from calcium phosphate coatings and its possible effect on enhanced bone integration', Acta Odontologica Scandinavica, 67, 333±340. Wieland M, Textor M, Chehroudi B, Brunette D M (2005), `Synergistic interaction of topographic features in the production of bone-like nodules on Ti surfaces by rat osteoblasts', Biomaterials, 26, 1119±1130. Wu S L, Liu X M, Hu T, Chu P K, Ho J P Y, Chan Y L, Yeung K W K, Chu C L, Hung T F, Huo K F, et al. (2008), `A biomimetic hierarchical scaffold: Natural growth of nanotitanates on three-dimensional microporous Ti-based metals', Nano Letters, 8, 3803±3808. Yan Y, Neville A, Dowson D (2006), `Understanding the role of corrosion in the degradation of metal-on-metal implants', Proceedings of the Institution of Mechanical Engineers Part H ± Journal of Engineering in Medicine, 220, 173±181. Yang L, Sheldon B W, Webster T J (2009a), `The impact of diamond nanocrystallinity on osteoblast functions', Biomaterials, 30, 3458±3465. Yang L, Sheldon B W, Webster T J (2009b), `Orthopedic nano diamond coatings: Control of surface properties and their impact on osteoblast adhesion and proliferation', Journal of Biomedical Materials Research Part A, 91A, 548±556. Yang Y Z, Kim K H, Ong J L (2005), `Review on calcium phosphate coatings produced using a sputtering process ± an alternative to plasma spraying', Biomaterials, 26, 327±337. Yildiz F, Yetim A F, Alsaran A, Efeoglu I (2009), `Wear and corrosion behaviour of various surface treated medical grade titanium alloy in bio-simulated environment', Wear, 267, 695±701. Yukna R A, Krauser J T, Callan D P (2000), `Multi-center clinical comparison of combination anorganic bovine-derived hydroxyapatite matrix (ABM)/cell-binding peptide (P-15) as a bone-replacement graft material in human periodontal osseus defects: 6 month results', Journal of Periodontology, 71, 1671±1679.

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Zhang L J, Webster T J (2009), `Nanotechnology and nanomaterials: Promises for improved tissue regeneration', Nano Today, 4, 66±80. Zhao G, Raines A L, Wieland M, Schwartz Z, Boyan B D (2007), `Requirement for both micron- and submicron scale structure for synergistic responses of osteoblasts to substrate surface energy and topography', Biomaterials, 28, 2821±2829. Zhao Q, Liu Y, Wang C, Wang S, Peng N, Jeynes C (2008), `Reduction of bacterial adhesion on ion-implanted stainless steel surfaces', Medical Engineering & Physics, 30, 341±349. Zinger O, Zhao G, Schwartz Z, Simpson J, Wieland M, Landolt D, Boyan B (2005), `Differential regulation of osteoblasts by substrate microstructural features', Biomaterials, 26, 1837±1847. Ziyad S, Haidar A, Reggie C, Hamdy A, Tabrizian M (2009), `Delivery of recombinant bone morphogenetic proteins for bone regeneration and repair. Part B: Delivery systems for bmps in orthopaedic and craniofacial tissue engineering', Biotechnology Letters, 31, 1825±1835.

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Index

-helices, 89 ABDMS see 4-aminobutyldimethylmonomethoxysilane acetylsalicylic acid, 25 acid etching, 124 ACL see anterior cruciate ligament actin cytoskeleton theory, 177 activated partial thromboplastin time, 65, 69 adenosine diphosphate, 60 AES see Auger Electron Spectroscopy AFM see atomic force microscopy air plasma spraying, 145 albumin, 59 Algipore, 381 alumina, 370 amino acids, 79, 83 chemical structures, 80 2-amino oleic acid, 51 (3-amino-propyl)trimethoxysilane, 47 4-aminobutyldimethylmonomethoxysilane, 43 ammonia plasma treatment, 287 anatase see titanium dioxide angioplasty, 257 angle resolved x-ray photoelectron spectroscopy, 215 anisotropic plasma etching, 4±5 anodic oxidation, 128 anodisation, 130, 132 anterior cruciate ligament, 314 anti-platelet drugs, 268 antibacterial molecules bacterial adherence and subsequent biofilm formation, Plate VI bacterial attachment on polyallylamine surfaces, Plate V covalent immobilisation, 294±304 established antibiotics, 296±8

furanones, 298±301 other cationic compounds, 296 quaternary ammonium compounds, 294±6 serrulatanes, 301±4 antibiotics, 296±8 9E10 antibody, 23 antithrombin III, 70 AOA see 2-amino oleic acid APS see air plasma spraying APTMS see (3-aminopropyl)trimethoxysilane APTT see activated partial thromboplastin time arginine-glycine-aspartic acid peptide, 348, 355, 359, 390 argon ion beam, 147 aromatic peptide amphiphiles, 91 arteriovenous graft placement, 257 artificial blood vessels, 187±8 artificial electrical prosthesis, 328 aseptic loosening, 369 aspirin see acetylsalicylic acid astrocytes, 325±7 atomic force microscopy, 238, 241 application, 241±3 contact mode image vs SEM image, 242 tapping mode data, 243 principle, 240 attachment point theory, 188 Auger Electron Spectroscopy, 208 Auger process, 208 autologous bypass conduits, 255 -hairpins, 89 -sheets, 89 -tricalcium phosphate, 372 Beer-Lambert law, 214

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402

Index

Bio-Oss, 376 bio-polymers, 335 bioactive agents, 267 bioactive coatings, 151 bioactives biomaterials surface modification to control infection, 284±305 antibacterial molecules covalent immobilisation, 294±304 future trends, 304±5 plasma-based strategies, 287±90 plasma polymers with metal nanoparticles or ions, 290±4 BioCoral, 376 biofilm, 286 biofunctionalisation, 261, 268±71, 333 covalently linked macromolecules, 270±1 passive functionalisation, 268±70 peptide capture, 271 techniques to enhance haemocompatibility, 263 BioHorizons, 388 biomaterials calcium phosphate deposition, 143±58 basic methods and applications, 144±55 future trends, 157±8 strengths and weaknesses, 156±7 energetics, 248±50 contact angle measurements, 249±50, 251±2 heparinisation, 56±74 bioactive molecule, 58 blood-biomaterial interaction, 59±62 future trends, 73±4 for improved blood compatibility, 62±73 nerve cell repair and regeneration, 325±38 future trends, 334±8 nervous system, 325±9 surface properties and effect on neural cells, 329±34 peptide functionalisation, 78±95 defining the biomaterial surface, 81 non-covalent peptide functionalisation by self-assembly, 88±91 peptide functionalised surfaces, 81±8 peptide functionality spatial control, 91±5 peptides and peptide functionalisation surfaces, 78±81

plasma polymerisation, 3±28 future trends, 26±8 intrinsic parameters, 13±17 overview, 4 plasma generation and system design, 4±8 plasma parameters, 8±13 potential biomaterials applications, 17±26 poly(ethylene glycol) covalent binding, 39±54 future trends, 52±4 principles and methods, 40±51 technologies and applications, 51±2 surface chemistry analysis techniques, 205±28 sample preparation and handling, 227±8 time of flight secondary ion mass spectrometry, 219±26 x-ray photoelectron spectroscopy, 207±19 surface modification for optimised interactions with blood, 255±73 biofunctionalisation, 268±71 physicochemical modification, 263±8 surface modification for optimised interactions with bone, 365±90 bone graft substitute materials, 375±82 joint replacements, 368±75 percutaneous bone-anchored devices, 382±90 surface modification for stem cell growth and differentiation control, 344±60 future trends, 359±60 surfaces for directing stem cell differentiation, 350±6 surfaces for maintaining the differentiated cell phenotype, 356±9 surfaces for stem cell expansion, 346±50 surface modification to optimise interactions with soft tissues, 309±22 kidneys, 312±14 liver, 309±12 skeletal muscle, 316±19 skin, 319±21 tendons/ligaments, 314±16 surface modification with bioactives to control infection, 284±305

ß Woodhead Publishing Limited, 2011

Index antibacterial molecules covalent immobilisation, 294±304 future trends, 304±5 plasma-based strategies, 287±90 plasma polymers with metal nanoparticles or ions, 290±4 surface morphology and topography, 233±45 confocal microscopy, 235±6, 237 profilometry, 243±5 scanning electron microscopy, 236±8 scanning force or atomic force microscopy, 238, 240±3 techniques for characterisation, 234 surface structure, morphology and topography analysis techniques, 232±53 future trends, 250, 253 surface structure and spatial distribution, 245±8 confocal Raman spectroscopy, 247±8 x-ray reflectivity and scattering, 245±7 surface topography to control cellular response, 169±91 current issues and future trends, 190±1 effect on cell behaviour, 175±82 manufacturing surface topography, 170±5 micron and nano definition, 170 photolithography and colloidal lithography, 171 technologies and potential applications, 182±6 tissue regeneration, 186±90 3i Biomet Nanotite implant, 387 biomimetic coatings, 154±5 blastocyst, 346 blood biofunctionalisation, 268±71 covalently linked macromolecules, 270±1 passive functionalisation, 268±70 peptide capture, 271 in situ graft endothelialisation, 272 techniques, 263 biomaterials surface modification for improved compatibility, 56±74 bioactive molecule, 58 blood-biomaterial interaction, 59±62 future trends, 73±4 for improved blood compatibility, 62±73

403

biomaterials surface modification for optimised interactions, 255±73 blood-biomaterial interface, 256±7 blood coagulation, 257 blood-contacting devices, 261±3 blood flow, 259 cell-material interactions, 259 cell-receptor mediated interactions, 259±60 endothelialisation, 260±1 thrombosis and thrombogenicity, 257±8 physicochemical modification, 263±8 bioactive agents, 267 immunosuppressant, cancer chemotherapy and anti-platelet drugs, 268 ordered patterning, 264±5 surface passivation and bulk chemical modification, 266±7 surface roughness, 263±4 UV light and plasma surface modification, 265±6 blood clotting time, 57 blood coagulation, 257 blood-contacting devices, 261±3 strategies for development, 262 blood flow, 259 BMP see bone morphogenetic proteins BMP-2 see bone morphogenetic proteins-2 BMP-4, 386 BMP-7 see bone morphogenetic proteins7 Boc see t-butoxycarbonyl Boc strategy, 83 Bohm sheath criterion, 17 bone biomaterials surface modification for optimised interactions, 365±90 bone-contacting materials surface modifications applications, 368 bone graft substitute materials, 375±82 joint replacements, 368±75 percutaneous bone-anchored devices, 382±90 human acetabular cup, 367 human femur showing trabecular and dense cortical bone, 366 bone graft commercial graft materials, 378 future developments, 381±2 controlling the rate of resorption, 381

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404

Index

improving the mechanical properties of graft materials, 381±2 surface modifications to combat infection, 382 materials grouping, 376±8 materials from animal bone, 376 materials from natural sources that resemble bone, 376 totally synthetic materials, 378 materials that mimic bone architecture, 377 maxillary sinus augmentation, 379 substitute materials graft topography and architecture, 379±80 organic surface modifications, 380±1 surface modification for optimised interactions, 375±82 bone grafting, 375 bone-like apatite, 154 bone morphogenetic proteins, 380 bone morphogenetic proteins-2, 380 bone morphogenetic proteins-7, 380 bottom-up methods, 81 bulk chemical modification, 266±7 calcium phosphate deposition basic methods and applications, 144±55 apatite formation, 156 biomimetic coatings, 154±6 electrochemical deposition, 151±2 electrophoretic deposition, 150±1 hot isostatic pressing, 153±4 ion beam assisted deposition, 147 magnetron sputtering deposition, 149±50 micro-arc oxidation, 148±9 plasma-sprayed HA coating on titanium alloy substrate, 146 pulsed laser physical vapour deposition, 148 sol-gel methods, 152±3 thermal spray deposition, 145±7 biomaterials surface modification, 143±58 future trends, 157±8 magnetron-sputtered HA coating on titanium alloy substrate, Plate II strengths and weaknesses, 156±7 calcium phosphates, 144, 148, 154 calcium sulphate hemihydrate, 381 calcium sulphate/PLLA composite, 381 cancer chemotherapy drugs, 268

carbon nanotubes, 337 carbonate-substituted apatite, 152 1,10 -carbonyldiimidazole, 45 carboxylic acid, 23, 48, 49 cardiopulmonary bypass circuit, 71±2 Carmeda BioActive Surface, 67, 68 PVC-DEHP surface modification, 68 casting, 331 cathodic polarisation, 128 CBAS see Carmeda BioActive Surface CDI see 1,10 -carbonyldiimidazole cellular response control by biomaterial surface topography, 169±91 mechanisms, 176±80, 181 cemented joints, 371 central nervous system, 325±6 Cerasorb, 378 Charnley stem, 372 chemical modification, 331 chemical shift effect, 207, 211 chemisorption, 123 chitosan, 319, 332 CIJ see Continuous Inkjet click chemistry, 53 CLSM see confocal laser scanning microscopy CNT see carbon nanotubes coagulation, 60 cobalt chromium molybdenum alloys, 370 collagen-like peptides, 94 collagen type I, 20, 354 collagen type II, 354 collagen type IV, 358 colloidal lithography, 171, 173 commercially pure titanium, 384 elastic modulus, 390 confocal laser scanning microscopy, 235±6 confocal microscopy, 235±6 application, 236 cancellous bone, 237 principle, 235 confocal Raman spectroscopy, 247±8 application, 248 long-term retrieved cup, wear index, nonwear and wear zone, Plate IV contact angle, 249 measurements, 249±50, 251±2 application, 250 surface roughness increase, 251±2 Young's force balance, 249 contamination, 188±9

ß Woodhead Publishing Limited, 2011

Index Continuous Inkjet, 93 copper, 293 covalent binding, 40, 51 biomaterials surface modification, 52±4 future trends, 52±4 principles and methods, 40±51 technologies and applications, 51±2 cpTi see commercially pure titanium cross linking, 369 Dacron, 270 delta-like-1, 23 dense cortical bone, 365 Density Functional Theory, 107 dental implantology, 186±7 dental root implants, 382±3 deposition, 134±5 device-related infections, 285±6 plasma-based strategies, 287±90 dextrans, 266±7 DFT see Density Functional Theory diamond-like coatings, 374 DII1 see delta-like-1 dip-pen nanolithography, 94 direct laser fabrication, 390 DLC see diamond-like coatings DOD see Drop-on-Demand doublets, 218±19 DPN see dip-pen nanolithography DRI see device-related infections Drop-on-Demand, 93 EGF see epidermal growth factor EILDVPST, 348 elastin-like polymers, 335 electrochemical deposition, 151±2 electrolytic deposition, 128 electron beam lithography, 172±3, 331 electron microscopy, 205 electron spectroscopy for chemical analysis, 67 see also x-ray photoelectron spectroscopy electrophoresis, 133 electrophoretic deposition, 128, 133±5, 150±1 electrospinning, 336 elemental quantification, 210±11 embossing, 331 embryonic stem cells, 183, 345±6 End-point attachment, 67 endothelial progenitor cells, 260 endothelialisation, 260±1 energetics, 248±50

405

contact angle measurements, 249±50, 251±2 EPC see endothelial progenitor cells EPD see electrophoretic deposition EPDIM, 319 epidermal growth factor, 348 ePTFE see extended polytetrafluoroethylene erythrocytes, 61 ESC see embryonic stem cells ESCA see electron spectroscopy for chemical analysis ethylene amine, 64 extended polytetrafluoroethylene, 265±6, 267 extracellular matrix molecules, 310 skeletal muscle, 317±19 skin, 319±20 fibrillin-1, 359 fibrinogen, 59, 70, 71, 242, 264 fibroblasts, 319 fibroin, 315 fibronectin, 316, 317, 347, 353, 359 fibronectin receptor, 353 filopodia, 178 fimbrolides, 298 fluorenylmethoxycarbonyl, 83, 91 fluoride, 152, 153 Fmoc see fluorenylmethoxycarbonyl Fmoc strategy, 83, 85 Freestyle Aortic Root Bioprosthesis, 52 Friadent plus, 386 FRIOS Algipore, 376 furanones, 298±301, 302, 303±4 attachment strategy for coated contact lenses, 301 chemical structures, 299 grafting scheme for covalent immobilisation, 300 Furlong stem, 372 gacyclidine, 327 galactose, 311 Gaseous Electronics Conference, 6 Gaussian-Lorentzian functions, 212 GBH see graphite-benzalkonium-heparin GEC see Gaseous Electronics Conference Genous R stent, 52 glutaraldehyde, 64 graft thrombogenicity, 256 grafting, 42 graphite-benzalkonium-heparin, 64 grazing incidence scattering, 247

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406

Index

haematopoietic stem/progenitor cells, 347 Hartree-Fock method, 107, 108, 109 heart valve replacement, 255±6 heparin, 24, 270, 311 anticoagulation effect, 58 chemical structure, 58 covalent binding, 64±6, 67 bound to collagen, 64 bound to polyurethane catheters, 64 graft copolymerisation on PU surface, 66 PEO grafting and heparin immobilisation, 67 immobilisation on polymer surface, 63 ionic binding, 63±4 physical blending for control release, 67±8 heparin immobilisation, 62 first method, 66 second method, 66 heparin-sirolimus, 268 heparinisation blood-biomaterial interaction, 59±62 coagulation, 60 coagulation system human proteins, 60 complement activation, 61 erythrocytes and leukocytes, 61 fibrinolysis, 61 interrelationships, 62 platelet reactions, 59±60 protein adsorption, 59 evaluation effect on blood compatibility, 69±73 clinical effect, 72 thrombus formation effect, 69±72 surface heparinisation, 63±8 Carmeda BioActive Surface, 67, 68 heparin covalent binding, 64±6, 67 heparin ionic binding, 63±4 heparin physical blending for control release, 67±8 surface modification for improved blood compatibility, 56±74 analysis of heparinisation degree, 68±9 bioactive molecule, 58 definition, 62 future trends, 73±4 heparin chemical structure, 58 heparin immobilisation on polymer surface, 63

material surface and blood compatibility co-relationship hypothesis, 57 methodological approaches, 62±8 heparinised surface, 62, 64 heparinoids, 58 hepatocyte, 309±10, 312 hESC see human embryonic stem cells hexamethylene diisocyanate, 41 high-energy photons, 16 high-velocity oxy-fuel, 146 hirudin, 270 HMDI see hexamethylene diisocyanate HNO3 solutions, 125 H2O2 surface treatment, 125 hot isostatic pressing, 153±4 Howell unit, 58 HSPC see haematopoietic stem/progenitor cells human embryonic stem cells, 346 HVOF see high-velocity oxy-fuel hydration, 111 hydroxyapatite, 144, 372±3, 375 hyper spectral imaging, 226 IKVAV peptides see isolucinelysinevaline-alanine-valine peptides IMFP see inelastic mean free paths Immix Extenders, 378 immunosuppressant drugs, 268 InductOs, 380 inelastic mean free paths, 214±15 infection bone graft materials, 382 control by biomaterials surface modification with bioactives, 284±305 antibacterial molecules covalent immobilisation, 294±304 future trends, 304±5 plasma-based strategies, 287±90 plasma polymers with metal nanoparticles or ions, 290±4 joint arthroplasty, 375 inflammation, 189±90 inkjet printing, 93±4 integrin receptors, 259±60 integrins, 349 intimal hyperplasia, 258 ion implantation, 387 IROX, 126 isocyanates, 41±2 isolucinelysine-valine-alanine-valine peptides, 333±4, 336, 354, 355, 359

ß Woodhead Publishing Limited, 2011

Index joint replacements articulating surfaces wear reduction methods, 369±71 cemented joints, 371 cementless hip replacements modular designs, 370 future developments, 373±5 modifications to improve fixation and resist infection, 374±5 wear resistance improvement, 373±4 non-cemented joints, 371±3 metal-on-metal hip-resurfacing device, 373 surface modification for optimised interactions, 368±75 kidneys, 312±14 KRSR, 390 lactose, 311 laminin, 319, 333, 347, 357±8 laser ablation, 238 laser holography, 172 Laser-Lok implant, 388, 389 laser scanning confocal microscopy see confocal laser scanning microscopy LCST see lower critical solution temperature LEED see low energy electron diffraction leukaemia inhibitor factor, 348 leukocytes, 61, 62 LIF see leukaemia inhibitor factor ligament reconstruction, 314 ligaments, 314±16 ligands, 259 lithography, 171 liver, 309±12 extracellular matrix molecules, 310 hepatocyte micrographs, 312 plasma treatment, 311 sugars, 311±12 low biofouling polymer coatings, 25±6 low energy electron diffraction, 109 low molecular weight heparin, 58 low pressure plasma spraying, 145 low temperature plasma, 5, 16 lower critical solution temperature, 349 LPPS see low pressure plasma spraying LQVQLSIR, 359 lysozyme, 296 magnetron sputtering deposition, 149±50 MALDI-ToF mass spectrometry, 92

407

MAO see micro-arc oxidation MAO-based treatments, 133 Matrigel, 354±5 mechanical interlocking, 372 MEMS see microelectromechanical systems technology mESC see murine embryonic stem cells metal alloys, 102, 105 metal oxides metallic biomaterials, 102±37 abbreviations and symbols, 142 future trends, 135±7 materials and methods for unpublished results, 141±2 native TiO2 chemical modification methods, 123±8 chemical etched titanium surface, 125 cleaning procedures, 123±4 sol-gel coatings, 126±8 wet chemical etching, 124 wet chemical oxidation, 125±6 native TiO2 electrochemical modifications, 128±35 anodic oxidation, 128 anodic polarisation and spark deposition, 130±3 cathodic polarisation, 128 electrochemical anodic and cathodic surface modifications, 129 electrophoretic deposition, 128, 133±5 enriched titanium oxide SEM and EDS analysis, 134 voltage/time evolution scheme, 131 surface oxides in metallic medical devices, 102±5 metal orthopaedic and dental prosthetic components, 103 orthopaedic metallic implant materials characteristics, 104 titanium dioxide surfaces, 107±21 adsorption phenomena, 117±21 hydration phenomena, 110±16 theoretical description, 107±10 titanium oxides technologies for tailoring on titanium, 121±35 titanium alloys, 106±7 on titanium implants, 105±21 metallic biomaterials, 102±37 methylprednisolone, 327 micro-arc oxidation, 148±9, 156±7 micro-patterning, 331±2

ß Woodhead Publishing Limited, 2011

408

Index

micro-topography, 351±2 microcontact printing, 93 microelectromechanical systems technology, 314 micron-scale, 170 molecular dynamics simulations, 111, 117 classical, 112 molecular self-assembly, 88 monomer deficient regime, 9 Moore's law, 172 mPEG see poly(ethylene glycol) monomethyl ether Mulliken charge analysis, 107 murine embryonic fibroblast feeder layer, 347 murine embryonic fibroblast (MEF) feeder layer, 347 murine embryonic stem cells, 346 MySkin, 18, 21, 320±1 N-isopropyl acrylamide, 25 Nafion, 20 Nanog, 183 nanoimprinting, 180 fibroblasts, 181 nanoscale, 170 nanotechnology, 329±30 nanotopography, 170±5, 352±3 natural film oxide, 103 near-field scanning optical microscopy, 250 nerve cells biomaterial surface modification for repair and regeneration, 325±38 future trends, 334±8 nervous system, 325±9 astrocytes in injury, 327 biomaterials used, 329 central nervous system, 325±6 characteristics and general behaviour after injury, 326±7 current therapies, 327±8 materials for neural regeneration, 328±9 surface properties and effect on neural cells, 329±34 glial cells and neurons co-culture, 330 peptide sequences for biomaterial functionalisation, 334 technologies and general applications, 330±2 topographical and chemical surface modifications, 332±4

nerve growth factor, 333 nerve regeneration, 186 nervous system, 325±9 astrocytes in injury, 327 central nervous system, 325±6 characteristics and general behaviour after injury, 326±7 current therapies, 327±8 materials for neural regeneration, 328±9 Nestin, 185 Neu5Gc, 346 neural regeneration, 328±9 biomaterials used, 329 future trends, 334±8 carbon nanotubes, 337 computer aided design scaffolds and solid free-form fabrication, 337±8 3D scaffolds, 334 elastin-like polymer electrospun fibres, 335 electrospinning, 336 genetically engineered bio-polymers, 335 peptide amphiphile, 335±6 self-assembling molecules, 334 neural stem cells, 185±6 neurons, 326 NGF see nerve growth factor NiPAAm see N-ispopropyl acrylamide Nobel Biocare MKIII, 386 Nobel Biocare TiUnite, 386 non-cemented joints, 371±3 non-surface damaging processes, 155 Norian SRS, 378 Novobiocin, 297±8 OCP see octacalcium phosphate octacalcium phosphate, 148 optical microscopy, 235 Osigraft, 380 osseointegration, 376 and dental implant, 384 influence of surface roughness and topography, 385±7 Dentsply Ankylos implant with Friadent Plus, 386 interactions of blood with rough and smooth dental implant surface, 385 sequence of events, 385 osteoblasts, 365 osteoclastogenesis, 375 osteoclasts, 365

ß Woodhead Publishing Limited, 2011

Index osteogenic differentiation, 356 OsteoGraf, 378 Osteograf /N, 381 osteoinduction, 379 osteolysis, 369 osteoprogenitor protein 1, 380 oxinium, 370 Paclitaxel, 268 pair distribution function, 115 passive functionalisation, 268±70 PCA see principal component analysis PCL see polycaprolactone PDF see pair distribution function PDMS see poly(dimethylsiloxane) PEG see poly(ethylene glycol) Pepgen P15, 380, 381 peptide amphiphile, 335±6 peptide amphiphiles, 89, 91 peptide capture, 271 peptide functionalisation biomaterials surface modification, 78±95 biomaterial surface defining, 81 non-covalent by self-assembly, 88±91 RADA, peptide amphiphiles and Fmoc-peptides self-assemble, 90 peptide functionalised surfaces, 81±8 common activating agents in SPPS, 86 covalent approach, 82 functional groups in peptides covalent attachment, 87±8 one-step functionalisation, 85±8 stepwise synthesis, 83±5 typical protecting groups for SPPS, 84 peptide functionality spatial control, 91±5 dip-pen nanolithography, 94 inkjet printing, 93±4 photolithography, 92 soft lithography, 93 SPOT-synthesis, 92±3 peptides and peptide functionalisation surfaces, 78±81 amino acids chemical structures, 80 percutaneous bone-anchored devices dental implants surfaces evolution, 383 interactions of blood with rough and smooth surfaces, 385 microstructured cpTi dental implant, 384

409

future developments, 388±90 biomimetic surfaces, 388±90 modifying stiffness of implants, 390 loss of crestal bone height and platform switching, 388, 389 BioHorizons Laser-Lok implant, 389 crestal bone loss, 388 surface modification for optimised interactions, 382±90 dental implant and osseointegration, 384 dental root implants, 382±3 ion implantation, 387 osseointegration sequence of events, 385 surface roughness and topography influence on osseointegration, 385±7 percutaneous transluminal coronary angioplasty, 255 Persantin, 268 PET see poly(ethylene terephthalate) PEU see poly(ether)urethane phase imaging, 242 photo-deprotection, 92 photoionisation cross section, 210 photolithography, 92, 170±2, 331 physisorption, 123 plasma, 4 density, 17 generation and system design, 4±8 parameters, 8±13 power, 9±12 pressure, 13 plasma coating technology, 370 plasma enhanced chemical vapour deposition, 293 plasma immersion ion implantation, 293 plasma polymerisation, 285, 320±1 biomaterials surface modification, 3±28 future trends, 26±8 acrylic acid plasma polymer deposit, 27 deposited microplasma polymerised styrene patterns, 28 intrinsic parameters, 13±17 m/z 73 peak ion energy, 15 plasma density and ion flux, 16±17 potentials in plasma and at surfaces, 13±16 temperature, 16 overview, 4 plasma generation and system design, 4±8

ß Woodhead Publishing Limited, 2011

410

Index

typical RF plasma system, 7 plasma parameters, 8±13 functional group retention in plasma polymer film, 12 monomers flow rate, 8±9 plasma power, 9±12 plasma pressure, 13 power/flow rate ratio effect on functional group retention, 10 plasma polymers, 19±26 amount of bound heparin, 24 bioactive incorporation, 24±5 cell attachment, 20±1 function and structure retention, 25±6 functional group retention, 23±4 keratinocytes attachment to plasmacopolymerised surfaces, 21 protein binding, 22±3 spatial patterning, 22 structural retention, 25 surface wettability, 19±20 potential biomaterials applications, 17±26 interaction at the biological± biomaterial interface, 19 rat bone cancer cells attachment to polystyrene surface, Plate I plasma polymers, 19±26, 287±90 with incorporated metal nanoparticles or ions, 290±4 other metallic ions, 293±4 silver, 290±3 with metal nanoparticles or ions, 290±4 other metallic ions, 293±4 silver, 290±3 silver release, 291 plasma spraying technique, 144, 145 plasma surface modification, 265±6 plasma treatment, 311 plasminogen, 62 Plaster of Paris see calcium sulphate hemihydrate platelet adhesion, 65, 70 platelets, 59±60 platform switching, 388 PLL-g-PEG see poly(L-lysine)-graftedpoly(ethylene glycol) Pluronic F127, 350 PMMA see polymethyl methacrylate PNIPAM see poly(Nisopropylacrylamide) poly(2-methacryloyloxyethyl phosphorylcholine), 374

poly-N-p-vinylbenzyl-D-lactonamide, 311 poly(-hydroxy esters), 346 polycaprolactone, 151, 155 polydimethylsiloxane, 242, 311 poly(dimethylsiloxane), 93 polyelectrolyte multilayer films, 269 poly(ether)urethane, 266 poly(ethylene glycol) covalent binding bifunctional PEGs, 43±4 heterobifunctional PEG chemical structure, 43 homobifunctional diamino-PEG, 44 surface activation using ABDMS and succininc anhydride, 44 biomaterials surface modification, 39±54 carboxylic acid derivatives, 47±8 covalent binding of mPEG via an acylation reaction, 48 mPEG-acid chloride derivative formation, 47 coupling compounds, 44±7 1,10 -carbonyldiimidazole, 45 carbonate formation promoted by CDI, 45 EDC, DIC and DCC, 44 ester formation promoted by CDI, 45 mPEG CDI activation, 46 N,N-disuccinimidyl carbonate, 46 PEI DSC coupling to surface hydroxyl groups, 46 epoxides, 49±51 electrophilic PEG coupled to nucleophilic surface, 50 homobifunctional diglycidyl-PEG ether, 49 nucleophilic PEG coupled to electrophilic surface, 50 PEG-epoxide preparation, 50 future trends, 52±4 biomolecule-PEG-cyclopentadiene conjugate, 53 biomolecules immobilisation, 53 isocyanates, 41±2 PU-PEG sheet coupled through HMDI linker, 41 TDI-modified PEG, 42 poly(ethylene glycol) monomethyl ether, 42±3 chemical structure, 42 principles and methods, 40±51 siloxane tethers, 48±9 PEG-silanes possessing siloxane tethers, 49

ß Woodhead Publishing Limited, 2011

Index trialkoxysilylated PEGs, 49 technologies and applications, 51±2 2-amino oleic acid, 51 poly(ethylene glycol) monomethyl ether, 42±3 poly(ethylene terephthalate), 238 poly(l-lactide-co--caprolactone), 313 poly(l-lysine)-grafted-poly(ethylene glycol), 389 polymer demixing, 173±5 polymethyl methacrylate, 332, 352 polymethylmethacrylate, 144 poly(N-isopropylacrylamide), 349 polytetrafluoroethylene, 266, 270 polyurethane, 67 power, 9 power deficient regime, 9 principal component analysis, 224 profilometry, 243±5 application, 244±5 SL Ti2 surfaces, 244 surface roughness changes monitoring, 244 ProOsteon, 376 propionaldehyde, 295 protein adsorption, 180±2 PSM see plasma surface modification PTCA see percutaneous transluminal coronary angioplasty PTFE see polytetrafluoroethylene pulsed laser physical vapour deposition, 148 QAC see quaternary ammonium compounds quantum DFT method, 112 quaternary ammonium compounds, 294±6 radio frequency, 5 radio frequency magnetron sputtering, 149 Raman spectroscopy, 247±8 recombinant human bone morphogenetic proteins, 380 reductive amination, 43 reflectometry, 245±6 Repifermin, 269 RF see radio frequency RGD peptide see arginine-glycineaspartic acid peptide roughening, 372 Rutherford backscattering, 132 rutile see titanium dioxide

411

SAM see self-assembled monolayers SAMs see self-assembled monolayers SBF see simulated body fluid scanning electron microscopy, 236±8, 239, 250, 253 application, 238 laser-ablated regions on PET surface, 239 schematic diagram, 237 scanning force microscopy, 238, 241 application, 241±3 Schwann cells, 186, 332 SCI see spinal cord injury secondary ion mass spectrometry, 205 self-assembled monolayers, 20, 81, 319± 20, 321, 334 self-bias potential, 14±15 SEM see scanning electron microscopy serecin, 315 serotonin, 60 serrulatanes, 301±4 skeleton and structures, 302 SFF see solid free-form fabrication SFM see scanning force microscopy shake-off, 218 shake-up satellites, 218 shear stress, 259 silane, 295 silicon fluoride, 375 silicon-HA, 153 silk, 314±16 silver, 290±3, 295±6 silver-hydroxyapatite nanocomposite, 375 SIMS see secondary ion mass spectrometry simulated body fluid, 128, 152, 154 Sirolimus, 268 skeletal muscle, 316±19 extracellular matrix molecules and their orientation, 317±19 tissue engineering, 318 topographical, 317 skeletal stem cells, 184±5 skin, 319±21 de-epidermalized dermis staining from keratinocyte transfer, Plate VII extracellular matrix proteins/peptides, 319±20 plasma polymerisation, 320±1 SLActive, 386 sodium cyanoborohydride, 42 sodium hydroxide treatment, 155 soft lithography, 93, 331

ß Woodhead Publishing Limited, 2011

412

Index

soft tissues biomaterials surface modification for optimised interactions, 309±22 kidneys, 312±14 liver, 309±12 skeletal muscle, 316±19 skin, 319±21 sol-gel coatings, 126±8 sol-gel methods, 152±3 solid free-form fabrication, 337±8 spinal cord injury, 326 SPOT-synthesis, 92±3 SPR see surface plasmon resonance static limit, 220 stem cells, 182±3 biomaterials surface modification for growth and differentiation control, 344±60 adipose-derived rat stem cells cultured on tissue culture plastic, 345 future trends, 359±60 NG108-15 neuronal cells cultured on fibronectin and laminin, 358 surfaces for maintaining the differentiated cell phenotype, 356±9 differentiation, 184 surfaces for directing stem cell differentiation, 350±6 growth factors, 356 peptides, 355 proteins, 353±5 surface chemistry modification, 351 surface topography modification, 351±3 surfaces for stem cell expansion, 346±50 enzyme-free subculture, 349±50 growth factors, 348±9 peptides, 348 proteins, 347 surface chemistry modification, 346±7 see also specific type Stepwise Surface process, 127 Straumann, 378, 386 strontium ions, 375 sub-micron, 170 sugars, 311±12 superlubricious coatings, 373±4 surface charge, 217 surface chemistry, 232

surface chemistry analysis techniques, 205±28 organisations, conferences and Internet resources, 228 sample preparation and handling, 227 time of flight secondary ion mass spectrometry, 219±26 image analysis, 226 principles, 219±21 spectrum interpretation, 222±6 x-ray photoelectron spectroscopy, 207±19 chemical state information, 211±14 depth information, 214±16 electron photoemission process, 207±10 elemental quantification, 210±11 imaging, 216±17 other special features, 217±19 surface energy, 248 surface modification antibacterial molecules covalent immobilisation, 294±304 established antibiotics, 296±8 furanones, 298±301 other cationic compounds, 296 quaternary ammonium compounds, 294±6 serrulatanes, 301±4 bioactives to control infection, 284±305 future trends, 304±5 plasma-based strategies, 287±90 biomaterials for kidneys, 312±14 poly(L-lactide-co--caprolactone) microspheres, 313 biomaterials for liver, 309±12 extracellular matrix molecules, 310 hepatocyte micrographs, 312 plasma treatment, 311 sugars, 311±12 biomaterials for skeletal muscle, 316±19 extracellular matrix molecules and their orientation, 317±19 tissue engineering, 318 topographical, 317 biomaterials for tendons/ligaments, 314±16 poly(lactic-co-glycolic acid) tissue engineered construct, 315 tenocytes micrographs, 316 biomaterials for the skin, 319±21

ß Woodhead Publishing Limited, 2011

Index extracellular matrix proteins/ peptides, 319±20 plasma polymerisation, 320±1 calcium phosphate deposition, 143±58 basic methods and applications, 144±55 future trends, 157±8 strengths and weaknesses, 156±7 heparinisation for improved blood compatibility, 56±74 bioactive molecule, 58 blood-biomaterial interaction, 59±62 future trends, 73±4 for improved blood compatibility, 62±73 nerve cell repair and regeneration, 325±38 future trends, 334±8 nervous system, 325±9 surface properties and effect on neural cells, 329±34 optimised interactions with blood, 255±73 biofunctionalisation, 268±71 physicochemical modification, 263±8 optimised interactions with bone, 365±90 bone graft substitute materials, 375±82 joint replacements, 368±75 percutaneous bone-anchored devices, 382±90 optimising interactions with soft tissues, 309±22 peptide functionalisation, 78±95 defining the biomaterial surface, 81 peptide functionalised surfaces, 81±8 peptides and peptide functionalisation surfaces, 78±81 plasma polymerisation, 3±28 plasma polymers with metal nanoparticles or ions, 290±4 other metallic ions, 293±4 silver, 290±3 silver release, 291 poly(ethylene glycol) covalent binding, 39±54 future trends, 52±4 principles and methods, 40±51 technologies and applications, 51±2 stem cell growth and differentiation control, 344±60 future trends, 359±60

413

surfaces for directing stem cell differentiation, 350±6 surfaces for maintaining the differentiated cell phenotype, 356±9 surfaces for stem cell expansion, 346±50 TiO2 modification, 122 surface morphology, 233±45 surface passivation, 266±7 surface plasmon resonance, 26 surface porosity, 264 surface properties, 329±34 surface roughness, 249, 263±4 surface tension, 248±9 surface topography, 233±45 control cellular response, 169±91 current issues and future trends, 190±1 micron and nano, 170 effect on cell behaviour, 175±82 chemistry vs topography, 182 filopodial interaction with nanosubstrates, 178 importance of feature size, 175±6 mechanisms of cellular response, 176±80, 181 nanoimprinting into fibroblasts, 181 nucleus morphology changes, 180 protein adsorption, 180±2 manufacture, 170±5 colloidal lithography, 173 electron-beam lithography, 172±3 laser holography, 172 photolithography, 170±2 polymer demixing, 173±5 technologies and potential applications, 182±6 embryonic stem cells, 183 isolation and expansion of cells in vitro, 182±3 neural stem cells, 185±6 skeletal stem cells, 184±5 stem cell differentiation, 184 tissue regeneration, 186±90 artificial blood vessel, 187±8 contamination, 188±9 dental implantology, 186±7 inflammatory response, 189±90 neural, 186 surface X-ray diffraction, 108 SurfaceSpectra version 4, 223 SXRD see surface X-ray diffraction Synthetic P15 peptide, 381

ß Woodhead Publishing Limited, 2011

414

Index

t-butoxycarbonyl, 83 TBI see traumatic brain injury TCPS see tissue culture polysterene TDI see toluene diisocyanate TDMAC see tridodecyl-methylammonium chloride Teflon-like plasma polymer films, 19±20 TEM see transmission electron microscopy tendon reconstruction, 314 tendons, 314±16 tenocytes micrographs, 316 tensegrity, 179 thermal spray deposition, 145±7 thrombogenicity, 258 thrombomodulin, 269 thrombosis, 257±8 thrombus formation, 60 effect, 69±72 time of flight secondary ion mass spectrometry, 219±26 image analysis, 226 tetraethylene glycol thiol on gold, 226 principles, 219±21 energy compensation in single stage mirror reflection, 221 primary ions and surface collision, 220 spectrum interpretation, 222±6 PCA score plot, 225 polylactic acid high mass resolution spectrum, 223 positive ion static SIMS spectrum, 222 predicted BAEC growth by partial least squares vs measured values, 225 predicted water contact angle from partial least squares model, Plate III TiO2 see titanium dioxide tissue culture polysterene, 20 tissue regeneration, 186±90 artificial blood vessel, 187±8 contamination, 188±9 dental implantology, 186±7 inflammatory response, 189±90 neural, 186 TissuePatchDural, 52 titanium alloys, 370±1 titanium dioxide, 121 chemical modification methods, 123±8 cleaning procedures, 123±4

sol-gel coatings, 126±8 wet chemical etching, 124 wet chemical oxidation, 125±6 electrochemical modification methods, 128±35 anodic oxidation, 128 anodic polarisation and spark deposition, 130±3 cathodic polarisation, 128 electrophoretic deposition, 128, 133±5 titanium dioxide surfaces, 107±21 adsorption phenomena, 117±21 albumin initial arrangements and fibronectin models, 119 proteins subdomain interaction energy, 119 hydration phenomena, 110±16 water molecules distribution, 115 water oxygens pair distribution function, 116 theoretical description, 107±10 most stable TiO2 surface, 107 titanium implants, 105±21 titanium oxides, 105±21 enriched titanium oxide, 134 Tixos, 390 ToF SIMS see time of flight secondary ion mass spectrometry toluene diisocyanate, 41 top-down methods, 81 transdifferentiation, 185 transmission electron microscopy, 238 transmission function, 210 traumatic brain injury, 326±7 tridodecyl-methylammonium chloride, 64 trypsin, 349 ultrahigh molecular weight polyethylene, 105, 369 ultraviolet light, 265±6 UV irradiation, 111 vacuum plasma spraying, 145 vascular graft endothelialisation, 261 Vaseline, 57 vibrational sum frequency spectroscopy, 116 vinyl acetate, 151±2 vitronectin, 347 von Willebrand factor, 71 VPS see vacuum plasma spraying Vroman effect, 59

ß Woodhead Publishing Limited, 2011

Index VSFS see vibrational sum frequency spectroscopy VUV see high-energy photons water contact angle measurement, 206±7, 224 Wenzel ratio, 250 wet chemical etching, 124 wet chemical oxidation, 125±6 work function, 207±8 X-ray photoelectron spectroscopy, 205, 207±19 chemical state information, 211±14 C1s binding energy for organic samples, 212 polylactic acid C1s score level, 213 depth information, 214±16 depth profiling, 216 inelastic mean free paths, 214±15 electron photoemission process, 207±10 atomic orbital nomenclature, 208 instrumentation, 209±10 photoelectron and Auger electron

415

emission processes, 208 XPS instrument, 209 elemental quantification, 210±11 poly(L-lactic acid) XPS survey scan, 211 imaging, 216±17 plasma polymerised allylamine strips, 217 other special features, 217±19 doublets, 218±19  bond shake up, 218 shake-up satellites, 218 spin-orbital coupling, 219 surface charge, 217 X-ray reflectivity, 111, 112, 245±7 applications, 247 sample geometry, 246 X-ray reflectometry, 246 X-ray scattering, 246±7 XPS see X-ray photoelectron spectroscopy YIGSR, 354, 355, 359

ß Woodhead Publishing Limited, 2011