Biomaterials and Tissue Engineering in Urology

Biomaterials and tissue engineering in urology

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Biomaterials and tissue engineering in urology Edited by John Denstedt and Anthony Atala

Oxford

Cambridge

New Delhi

Published by Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, UK www.woodheadpublishing.com Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi – 110002, India Published in North America by CRC Press LLC, 6000 Broken Sound Parkway, NW, Suite 300, Boca Raton, FL 33487, USA First published 2009, Woodhead Publishing Limited and CRC Press LLC © 2009, Woodhead Publishing Limited 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 publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from 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. Library of Congress Cataloging in Publication Data A catalog record for this book is available from the Library of Congress. Woodhead Publishing ISBN 978-1-84569-402-9 (book) Woodhead Publishing ISBN 978-1-84569-637-5 (e-book) CRC Press ISBN 978-1-4398-0177-2 CRC Press order number: N10041 The publishers’ 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 acid-free and elemental chlorine-free practices. Furthermore, the publishers ensure that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by SNP Best-set Typesetter Ltd., Hong Kong

Printed by TJ international Limited, Padstow, Cornwall, UK

Contents

Contributor contact details Preface

xv xxiii

Part I Fundamentals

1

1

3

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 1.10 1.11 1.12 2

2.1

Introduction to biofilms in urology P. Cadieux and G. Wignall, University of Western Ontario, Canada; R. Carriveau, University of Windsor Canada Introduction What is a biofilm and why do they form? Biofilm formation and structure Biofilms in general medicine Biofilms in urology Biofilm shedding and migration: infection spread and recurrence Resistance to host factors and antibiotics Current and future biofilm prevention and treatment strategies Future trends Conclusions Sources of further information and advice References In vivo models for ureteral stents M. K. Louie, A. J. Gamboa and R. V. Clayman, UCI Medical Center, USA Introduction

3 4 5 12 14 22 23 25 30 31 31 32 42

42 v

vi

Contents

2.2 2.3 2.4

Commonly used animal models Conclusion and future trends References

3

Models for the assessment of biofilm and encrustation formation on urological materials B. F. Gilmore, D. S. Jones and S. P. Gorman, Queen’s University Belfast, Northern Ireland; and H. Ceri, University of Calgary, Canada Introduction Development of urinary encrustation Assessment of biomaterial encrustation – in vitro models Dynamic flow-through models Batch flow or ‘static’ models Dynamic continuous flow models The MBEC-BESTTM assay Conclusions References

3.1 3.2 3.3 3.4 3.5 3.6 3.7 3.8 3.9

45 54 55

59

59 62 65 66 71 73 76 77 78

Part II Materials and design of urological devices

83

4

85

4.1 4.2 4.3 4.4 4.5 4.6 4.7 5

5.1 5.2 5.3 5.4 5.5 5.6 5.7

Ureteral stents: design and materials D. Lange and B. H. Chew, University of British Columbia, Canada Introduction Current stent biomaterials Stent coatings Stent design Drug-eluting stents Conclusions and future trends References Metal stents in the upper urinary tract E. Liatsikos, D. Karnabatidis, P. Kallidonis and D. Siablis, University of Patras, Greece Introduction Types of metal stents in the upper urinary tract Applications of metal stents Insertion techniques Complications and problems Virtual endoscopy and metal stents Extra-urinary drainage of the upper urinary tract

85 88 93 95 97 98 99 104

104 105 109 118 121 123 125

Contents

vii

5.8 5.9

Future trends References

126 129

6

Coated ureteral stents F. Cauda, Ospedale Koelliker, Italy; V. Cauda, Ludwig-Maximilians-Universität München, Germany; and C. Fiori, Azienda Ospedaliero-Universitaria San Luigi Gonzaga, Italy Introduction Methods Results Discussion Conclusions Acknowledgement References

134

6.1 6.2 6.3 6.4 6.5 6.6 6.7 7

7.1 7.2 7.3 7.4 7.5 7.6 7.7 7.8 7.9 7.10 7.11 7.12 8

8.1 8.2 8.3 8.4 8.5 8.6

Proteus mirabilis biofilm formation and catheter design D. J. Stickler, Cardiff University, UK Introduction Virulence factors Epidemiology of Proteus mirabilis infections The process of crystalline biofilm formation on catheters Antimicrobials in the prevention of catheter encrustation Factors that modulate the rate of Proteus mirabilis biofilm formation on catheters Urease inhibitors Catheter design Future trends Conclusions Sources of further information and advice References Self-lubricating catheter materials A. D. Woolfson, R. K. Malcolm, S. P. Gorman and S. D. McCullagh, Queen’s University Belfast, UK Introduction Silicone chemistry Self-lubricating silicone biomaterials Performance characteristics of self-lubricating silicone biomaterials Bioactive lubricious silicones Biomimetic lubricious silicones

134 137 139 151 154 154 155

157 157 158 162 163 170 175 178 179 183 184 185 185 191

191 192 197 199 201 203

viii

Contents

8.7 8.8 8.9

Toxicity and regulatory issues Conclusions References

203 205 206

9

Temporary urethral stents T. Tammela, Tampere University Hospital, Finland Introduction Indications for the use of stents Non-degradable temporary urethral stents Biodegradable urethral stents Future trends References

208

Penile implants G. Brock, University of Western Ontario, Canada Introduction Historical aspects of penile prosthesis development Biomaterials in current use Device infection Erosion resistance Summary Future trends References

226

9.1 9.2 9.3 9.4 9.5 9.6 10 10.1 10.2 10.3 10.4 10.5 10.6 10.7 10.8

208 209 211 214 221 222

226 227 230 232 234 235 236 237

Part III Urological tissue engineering

241

11

Artificial biomaterials for urological tissue engineering W. A. Farhat, The Hospital for Sick Children, Canada; and P. J. Geutjes, Radboud University Nijmegen Medical Centre, The Netherlands Introduction History of synthetic biomaterials used in urology Synthetic scaffolds Smart biomaterials Future trends References

243

Natural biomaterials for urological tissue engineering C. C. Roth and B. P. Kropp, The University of Oklahoma Health Sciences Center, USA; and E. Y. Cheng, Children’s Memorial Hospital in Chicago, USA Introduction

255

11.1 11.2 11.3 11.4 11.5 11.6 12

12.1

243 244 245 247 251 252

255

Contents

ix

12.2 12.3 12.4 12.5 12.6 12.7

Historical application of natural biomaterials Fundamental biomaterials Collagen-based extracellular matrices Future trends Sources of further information and advice References

256 257 261 274 275 275

13

Nanotechnology and urological tissue engineering B. S. Harrison and C. L. Ward, Wake Forest Institute for Regenerative Medicine, USA Introduction Rationale for nanomaterials in engineering tissue Use of nanomaterials as biomaterials Use of nanomaterials for aiding cell tracking Use of nanomaterials to improve drug delivery Conclusions Future trends Source of further information and advice References

281

13.1 13.2 13.3 13.4 13.5 13.6 13.7 13.8 13.9 14

14.1 14.2 14.3 14.4 14.5 14.6 15

15.1 15.2 15.3 15.4 15.5 15.6

Assessing the performance of tissue-engineered urological implants G. J. Christ, D. Burmeister, S. Vishwajit, Y. Jarajapu and K.-E. Andersson, Wake Forest Institute for Regenerative Medicine, USA Introduction The bladder Evaluation of engineered or regenerating tissues in vitro Bladder tissue engineering and regeneration Conclusions and future trends References Regenerative pharmacology and bladder regeneration K.-E. Andersson and G. J. Christ, Wake Forest Institute for Regenerative Medicine, USA Introduction Endogenous bladder regeneration Construction of a tissue or organ Development of an engineered bladder Implantation of the bladder construct in preclinical studies Preliminary clinical experience with neobladders

281 282 283 287 290 292 292 294 294

299

299 301 304 310 314 315

322

322 324 327 328 330 331

x

Contents

15.7 15.8 15.9

Conclusions Acknowledgement References

332 332 332

16

Autologous cell sources for urological applications Y. Zhang, Wake Forest Institute for Regenerative Medicine, USA Introduction Fully differentiated cells for urological reconstruction Stem/progenitor cells for urological reconstruction Cell tracking technology Conclusions Acknowledgements References

334

16.1 16.2 16.3 16.4 16.5 16.6 16.7 17

17.1 17.2 17.3 17.4 17.5 17.6 17.7 17.8 17.9 18

18.1 18.2 18.3 18.4 18.5

Embryonic stem cells, nuclear transfer and parthenogenesis-derived stem cells for urological reconstruction R. Dorin, J. Yamzon and C. J. Koh, Childrens Hospital Los Angeles and University of Southern California Keck School of Medicine, USA Introduction Principles of tissue engineering Stem cells: overview Embryonic stem cells Nuclear transfer Parthenogenesis Induced pluripotent stem cells Conclusions and future trends References Amniotic fluid and placental stem cells as a source for urological regenerative medicine P. De Coppi, UCL Institute of Child Health and Great Ormond Street Hospital, UK and Azienda Ospedaliera Università di Padova, Italy; G. Bartsch, University of Ulm, Germany; and A. Atala, Wake Forest Institute for Regenerative Medicine, USA Introduction Amniocentesis Differentiated cells from amniotic fluid Mesenchymal stem cells from amniotic fluid Amniotic fluid-derived stem cells

334 337 342 348 351 351 351

357

357 358 360 361 363 367 369 370 371

378

378 378 379 379 382

Contents

xi

18.6 18.7

Conclusions References

390 390

19

The use of adipose progenitor cells in urology D. S. Davé and L. V. Rodri´ guez, University of California Los Angeles, USA Introduction Nomenclature and origin of adipose progenitor cells Isolation procedures Molecular characterization Differentiation capacity of adipose-derived stem cells Applications in the field of urology Future trends References

395

19.1 19.2 19.3 19.4 19.5 19.6 19.7 19.8 20

20.1 20.2 20.3 20.4 20.5 20.6 20.7 20.8 20.9 21

21.1 21.2 21.3 21.4 21.5 21.6

Regenerative medicine of the urinary sphincter via an endoscopic approach M. C. Smaldone, University of Pittsburgh School of Medicine, USA; and M. B. Chancellor, William Beaumont Hospital, USA Introduction Neurophysiology of stress urinary incontinence Stem cell source for the injection therapy of stress urinary incontinence Role of muscle-derived stem cells in the delivery of neurotrophic factors Injection technique Current results of clinical studies Conclusions Acknowledgements References Regenerative medicine of the urinary sphincter via direct injection R. Yiou, CHU Henri Mondor, France Introduction Challenges with muscle precursor cell transfer The direct myofiber implantation procedure Direct injection of muscle precursor cells using minced muscle Conclusions and future trends References

395 397 397 399 402 404 411 412

422

422 424 427 433 434 436 438 438 439

445 445 446 447 450 451 451

xii

Contents

22

Regenerative medicine for the urethra T. Aboushwareb and A. Atala, Wake Forest Institute for Regenerative Medicine, USA; and A. Elkassaby, Ain Shams University, Egypt Introduction Synthetic scaffolds Biological (natural) polymers Conclusions Acknowledgement References

454

Penile reconstruction H.-J. Wang and J. J. Yoo, Wake Forest Institute for Regenerative Medicine, USA Introduction Basic principles of penile tissue engineering Engineering of functional corporal tissue Engineered penile prosthesis Reconstruction of the tunica albuginea Summary and future trends Acknowledgement References

470

Tissue engineering in reproductive medicine A. Sophonsritsuk and C. E. Bishop, Wake Forest Institute for Regenerative Medicine, USA Tissue engineering of the vagina Methods of vaginal tissue reconstitution Tissue engineering of the uterus Methods of uterine tissue reconstitution Tissue engineering of the ovarian tissue Method for culturing follicles Conclusions Acknowledgements References

482

Regenerative medicine of the kidney N. Guimaraes-Souza, R. Soler and J. J. Yoo, Wake Forest Institute for Regenerative Medicine, USA Introduction Basic components of renal tissue engineering Approaches for the regeneration of renal tissue Cell-based therapy for kidney disease

502

22.1 22.2 22.3 22.4 22.5 22.6 23

23.1 23.2 23.3 23.4 23.5 23.6 23.7 23.8 24

24.1 24.2 24.3 24.4 24.5 24.6 24.7 24.8 24.9 25

25.1 25.2 25.3 25.4

454 455 456 464 465 465

470 471 473 476 478 478 479 479

482 483 486 487 491 495 496 497 497

502 503 505 511

Contents

xiii

25.5 25.6 25.7

Summary Acknowledgement References

512 513 513

26

Stem cells and kidney regeneration S. Sedrakyan, L. Perin and R. E. De Filippo, University of Southern California, USA Introduction Endogenous stem cells Exogenous stem cells Conclusions References

518

Techniques for engineering bladder tissue A. Atala, Wake Forest Institute for Regenerative Medicine, USA Introduction Cells used in tissue engineering Biomaterials used in tissue engineering Bladder repair and replacement: current and future technologies Summary and conclusions References

532

Index

550

26.1 26.2 26.3 26.4 26.5 27

27.1 27.2 27.3 27.4 27.5 27.6

518 519 521 526 527

532 534 539 541 545 545

Contributor contact details

(* = main contact)

Editors

Chapter 1

Prof. John Denstedt University of Western Ontario 268 Grosvenor Street London Ontario N6A 4V2 Canada E-mail: John.Denstedt@sjhc. london.on.ca

Dr P. Cadieux* and Dr G. Wignall University of Western Ontario 268 Grosvenor Street London Ontario N6A 4V2 Canada E-mail: [email protected]

Prof. Anthony Atala, MD Department of Urology and Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

Dr R. Carriveau University of Windsor 401 Sunset Avenue Windsor Ontario N9B 3P4 Canada E-mail: [email protected]

Chapter 2 Michael K. Louie, Aldrin J. Gamboa and Dr Ralph V. Clayman* UCI Medical Center Department of Urology 333 City Boulevard West Suite 2100, Rt.81 Orange CA 92868 USA E-mail: [email protected]

xv

xvi

Contributor contact details

Chapter 3

Chapter 5

Dr Brendan F. Gilmore, Prof. David S. Jones and Prof. Sean P. Gorman* The School of Pharmacy Queen’s University Belfast 97 Lisburn Road Belfast BT9 7BL UK E-mail: [email protected]; [email protected]

Evangelos Liatsikos, MD, PhD* Assistant Professor Department of Urology University of Patras Greece E-mail: [email protected]

Prof. Howard Ceri Biofilm Research Group Department of Biological Sciences University of Calgary 2500 University Drive N.W. Calgary AB T2N 1N4 Canada

Dimitrios Karnabatidis, MD, PhD Assistant Professor Department of Radiology University of Patras Greece E-mail: [email protected] Panagiotis Kallidonis, MD Resident in training Department of Urology University of Patras Greece E-mail: [email protected]

Chapter 4 Dirk Lange, PhD and Ben H. Chew, MD* University of British Columbia Level 6-2775 Laurel St Vancouver BC V5Z 1M9 Canada E-mail: [email protected]

Prof. Dimitrios Siablis, MD, PhD Department of Radiology University of Patras Greece E-mail: [email protected]

Contributor contact details

Chapter 6

Chapter 8

Prof. Furio Cauda, MD Divisione di Urologia Ospedale Koelliker Corso G. Ferraris, 251-255 10134 Torino Italy E-mail: [email protected]

Prof. A. David Woolfson,* Dr R. Karl Malcolm, Prof. Sean P. Gorman and Dr Stephen D. McCullagh School of Pharmacy Queen’s University Belfast Belfast BT9 7BL UK E-mail: [email protected]

Dr Valentina Cauda, PhD* Department für Chemie und Biochemie Ludwig-Maximilians-Universität München Butenandtstr. 11-E 81377 München Germany E-mail: [email protected] Dott. Cristian Fiori, MD Division of Urology University of Turin ‘San Luigi Gonzaga’ Hospital Regione Gonzole 10-10043-Orbassano (Torino) Italy

Chapter 7 Dr David J. Stickler Cardiff School of Biosciences Cardiff University Cardiff CF10 3TL Wales UK E-mail: [email protected]

Chapter 9 Teuvo Tammela, MD, PhD Professor of Urology Department of Urology Tampere University Hospital Teiskontie 35/P.O. Box 2000 FIN-33521 Tampere Finland E-mail: [email protected]

Chapter 10 Gerald Brock Professor of Surgery University of Western Ontario 268 Grosvenor Street London Ontario N6A 4V2 Canada E-mail: [email protected]

xvii

xviii

Contributor contact details

Chapter 11

Chapter 12

Dr Walid A Farhat* The Hospital for Sick Children 555 University Avenue, RM 292 Toronto Ontario M5G 1X8 Canada E-mail: [email protected]

Christopher C. Roth, MD and Bradley P. Kropp, MD, FAAP* Pediatric Urology Children’s Hospital of Oklahoma The University of Oklahoma Health Sciences Center 920 Stanton L. Young Blvd WP 3150 Oklahoma City OK 73104 USA E-mail: [email protected]; [email protected]

Paul J. Geutjes 659 Urology Nijmegen Centre for Molecular Life Sciences (NCMLS) Radboud University Nijmegen Medical Centre Geert Grooteplein 26/28 6525 GA Nijmegen The Netherlands E-mail: [email protected]

Earl Y. Cheng, MD Associate Professor of Urology Children’s Memorial Hospital and The Feinberg School of Medicine at Northwestern University 2300 Children’s Plaza Box ⱅ24 Chicago IL 60614 USA E-mail: echeng@childrensmemorial. org

Chapter 13 Dr Benjamin S. Harrison* and Catherine L. Ward Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]; [email protected]

Contributor contact details

xix

Chapter 14

Chapter 17

G. J. Christ, PhD,* D. Burmeister, S. Vishwajit, Y. Jarajapu and K.-E. Andersson Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

Ryan Dorin, MD, Jonathan Yamzon, MD and Chester J. Koh, MD* Childrens Hospital Los Angeles 4650 Sunset Boulevard Mailstop 114 Los Angeles CA 90027 USA E-mail: [email protected]

Chapter 15

Prof. Paolo De Coppi, MD, PhD* Surgery Unit UCL Institute of Child Health and Great Ormond Street Hospital London UK and Stem Cell Processing Laboratory Department of Paediatrics Azienda Ospedaliera Università di Padova Italy E-mail: [email protected]

K.-E. Andersson, MD, PhD and George J. Christ, PhD* Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston Salem NC 27157-1094 USA E-mail: [email protected]

Chapter 16 Yuanyuan Zhang, MD, PhD Department of Urology and Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston Salem NC 27157-1094 USA E-mail: [email protected]

Chapter 18

Dr Georg Bartsch, Jr, MD Department of Urology University of Ulm Prittwitzstrasse 43 89075 Ulm Germany E-mail: [email protected] Prof. Anthony Atala, MD Department of Urology and Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

xx

Contributor contact details

Chapter 19

Chapter 21

Dhiren S. Davé, MD and Larissa V. Rodríguez, MD* Department of Urology David Geffen School of Medicine University of California Los Angeles USA E-mail: [email protected]. edu; [email protected]

René Yiou, MD, PhD MCU-PH AP-HP Service d’Urologie CHU Henri Mondor 51 avenue du Maréchal de Lattre de Tassigny 94010 Créteil France E-mail: [email protected] and [email protected]

Chapter 20 Marc C. Smaldone, MD* Department of Urology University of Pittsburgh School of Medicine Pittsburgh Pennsylvania USA E-mail: [email protected] Michael B. Chancellor, MD Director of Neurourology Research Department of Urology William Beaumont Hospital Royal Oak Michigan USA E-mail: [email protected]

Chapter 22 Tamer Aboushwareb* Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston Salem NC 27157-1094 USA E-mail: [email protected] Prof. Anthony Atala, MD Department of Urology and Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA Email: [email protected] Abdelwahab ElKassaby Professor of Urology Ain Shams University Cairo Egypt E-mail: [email protected]

Contributor contact details

xxi

Chapter 23

Chapter 26

Hung-Jen Wang, MD and James J. Yoo, MD, PhD* Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

Sargis Sedrakyan, Laura Perin and Dr Roger E. De Filippo* Childrens Hospital Los Angeles Division of Urology Saban Research Institute Developmental Biology Program Keck School of Medicine University of Southern California USA E-mail: [email protected]

Chapter 27 Chapter 24 Areepan Sophonsritsuk and Colin E. Bishop, PhD* Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

Chapter 25 Nadia Guimaraes-Souza, Roberto Soler, PhD and Dr James J. Yoo, MD, PhD* Department of Urology and Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

Prof. Anthony Atala, MD Department of Urology and Wake Forest Institute for Regenerative Medicine Wake Forest University School of Medicine Medical Center Boulevard Winston-Salem NC 27157-1094 USA E-mail: [email protected]

To my wife Carolyn, and our two daughters, Emily and Ellen J. D. To my wife Katherine, and our two children, Christopher and Zachary A. A.

Preface

One of the major advances of modern medicine has been the development of alloplastic devices designed either to replace human organs or to reestablish their functionality. The use of such devices, temporarily or permanently, has become so common in various fields of medicine, that it has been predicted that, ultimately, almost every human in technologically advanced societies will host a biomaterial. Biomaterials have been used in the urinary tract for many centuries and they have evolved from primitive metal tubes to complex coated polymeric materials crafted into various devices to provide drainage or improve the functionality of the native urinary tract structures. Urologists are amongst the most common implanters of biomaterials into the human body, amongst all of the subspecialties in surgery. Common biomaterial-based devices used in the urinary tract include urethral and suprapubic catheters, urethral and ureteral stents, nephrostomy tubes and devices such as urethral and penile implants. In recent years, different raw biomaterials have successfully made the transition from laboratory bench research to widespread use in the clinical realm. However, despite ongoing intensive basic science and clinical research, the ideal biomaterial substance has yet to be discovered. Nonetheless, there have been many significant advances in this field over the past several years and these are highlighted in the comprehensive contributions from the internationally based scientists and clinicians who have contributed chapters to this unique book focused on biomaterials and tissue engineering in the urinary tract. The compilation of chapters is a unique contribution to the field of urology and represents an up-to-date analysis of the spectrum of biomaterial device research and clinical application in the field. The widespread use of biomaterials in the urinary tract, such as in catheters and stents, will undoubtedly continue into the future. Ongoing research is essential to optimize biocompatibility and decrease biomaterial xxiii

xxiv

Preface

related complications, such as patient discomfort, bladder irritability, infection and encrustation of devices placed within the urinary tract. Many advances as outlined in this book have been made in recent years, including the development of new biocompatibility testing approaches, the introduction of new biomaterials and biomaterial coatings, the design of novel ureteral catheters and stents, and the promise of drug elution. With continued insights into biomaterial science we can expect further improvements in clinically available biomaterials. As this occurs, we will reach the ultimate goal, namely to develop the ideal biomaterial, which will result in a significant decrease in or perhaps even elimination of biomaterial-related morbidity. I thank all of the authors for their outstanding contributions to this book which truly represents a state-of-the art update on basic science developments and clinical use of biomaterial devices for use in the urinary tract.

1 Introduction to biofilms in urology P. C A D I E U X and G. W I G N A L L , University of Western Ontario, Canada; R. C A R R I V E AU, University of Windsor, Canada

Abstract: Despite significant resources and several decades of research aimed at their prevention and treatment, biofilm-associated infections continue to be the major cause of urological device failure. Numerous strategies have been targeted towards improving device design, biomaterial composition, surface characteristics and drug elution, but have been largely thwarted by microorganisms and their arsenal of attachment, host evasion, antimicrobial resistance and dissemination strategies. This is not entirely surprising considering that natural biofilm formation has been occurring for billions of years and remains a significant element of microbe survival and evolution. Thus, the fact that biofilms develop on and in the biomaterials and tissues of humans is largely an extension of this natural tendency and underscores why they are so difficult to combat. Biofilm structure and composition intrinsically offer a protective environment for microorganisms, shielding them from the shear stress of urine flow, attack by the host immune system and antimicrobials. Furthermore, many biofilm organisms go into a metabolically quiescent state that renders them more tolerant to antibiotics and host immune factors able to penetrate the biofilm matrix. Finally, most organisms causing biofilm-associated urinary tract infections originate from the host’s own oral cavity, skin, gastrointestinal and urogenital tracts and therefore have already adapted to many host defense mechanisms. Ultimately, while biofilms continue to hold the upper hand with respect to recurrent infections and urinary tract biomaterial use, significant progress has been made in understanding these dynamic microbial communities and novel approaches offer promise for biofilm prevention and removal. These include novel device designs, antimicrobials, anti-adhesive coatings, biodegradable polymers and biofilm-disrupting compounds and therapies. Key words: biofilm, antibiotics, urinary tract infection, ureteral stent, urethral catheter, uropathogen, conditioning film, antibiotic resistance.

1.1

Introduction

Biofilm-related infections remain the major cause of urological device failure despite millions of dollars and several decades of research targeted at their prevention and eradication. The plethora of strategies aimed at 3

4

Biomaterials and tissue engineering in urology

improving device design, biomaterial composition, surface characteristics and drug delivery have been generally thwarted by microbes and their abundance of attachment, host evasion, antimicrobial resistance, survival and dissemination strategies. It is important to acknowledge that the formation of biofilms by microorganisms in nature has been going on for billions of years and remains a major part of their current survival and evolution on the planet. Thus, the fact that biofilms develop on and in the tissues and biomaterials of humans is simply an extension of this natural tendency and largely explains why they are so difficult for us to combat. Furthermore, the majority of organisms that cause biofilm-related urinary tract infections originate from our own skin, oral cavity and gastrointestinal tract, and thus have already adapted to many of our defenses. In this chapter, we will first discuss biofilms at their most basic level including why and how they form. It is essential to understand the purpose, formation and structure of both natural and medical biofilms if we are to tackle the clinical problems associated with them and attempt to prevent and eradicate them from patients. Biofilms in medicine in general will follow and then transition into our focus of biofilms in urology and their clinical relevance. We will discuss the basic hydrodynamics of the upper and lower urinary tract along with how biofilms affect, and are affected by, urinary flow. Prosthetic versus non-prosthetic infections will be described as well as the ability of biofilms to resist antimicrobials, host urinary tract defenses and the immune system. The development and implications of chronic infections will be commented on and the effects that biofilms and biofilm-related encrustation have on them. This will lead into current treatment and prevention strategies including novel biomaterial surfaces and design, coatings, antimicrobial strategies and the promising field of biofilm-disrupting agents. We argue that although microbial biofilms continue to hold the upper hand in terms of recurrent infections and the use of biomaterials within the urinary tract, significant progress has been made in understanding these microbial communities and new strategies offer promise in the field.

1.2

What are biofilms and why do they form?

A biofilm can be defined as a community of microorganisms (bacterial, fungal, algal) attached to the interface of a liquid and a surface, and enveloped within a matrix of exopolysaccharides and other biological constituents (Costerton, 2007a; Jass et al., 2003). In the simplest terms, a biofilm is merely a mechanism used by microorganisms to remain in a favorable environment and promote their survival and reproduction within that environment. Thus, on a very basic level, a urinary biofilm on a ureteral stent is no different than those of the thermophilic bacterium Thermus aquaticus at a deep sea hydrothermal vent or Serratia marcescens on your shower wall (Langsrud et al., 2003; Stramer and Starzyk, 1978). In general, if a micro-

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organism comes in contact with a surface and successfully attaches to it, and the right growth conditions exist (temperature, nutrients, pH, osmolarity, etc.), it will adhere more securely and begin multiplying and forming a biofilm. In addition, biofilms provide a degree of protection for a microbial population. Although microorganisms can survive and multiply very efficiently as single cells within a liquid milieu (such as uropathogens in urine), they are completely exposed to any detrimental environmental conditions that might arise (i.e. increased temperature, host factors, antibiotics) and may be swept away from a favorable environment via shear forces within the surrounding fluid (such as during micturition). Thus, if a single-celled (planktonic) population is exposed to a sudden lethal change in temperature or high concentration of a toxic compound, all cells in the population may be killed. The structure and makeup of a biofilm protects many cells within the population by physically shielding them from the surrounding environment and inducing changes in gene expression that result in metabolic dormancy and/or increased resistance to many inhibitory substances (Donlan, 2003; Fujiwara et al., 1998). Although biofilms are far more complex in reality than described here, it is obvious that their formation improves microorganism population survival, growth and reproduction through multiple mechanisms. It is therefore not surprising that biofilms are relatively ubiquitous throughout every natural environment on the planet, ‘from tropical leaves to desert bolders’ (Costerton, 2007a), or more relevant here, from the vaginal epithelia to urinary tract devices. Clearly, biofilms are the greatest challenge faced by medicine and industry in terms of treating wounds, using biomaterials and designing compounds for the prevention and treatment of infection.

1.3

Biofilm formation and structure

At a macroscopic level, most biofilms appear fairly simple in structure, as uncomplicated clusters, chains and/or layers of organisms haphazardly attached to a surface and intermixed with extracellular slime. Based upon this view, one might consider biofilm development simply as the repeated process of single organisms from the surrounding fluid adhering themselves to an object with little or no purpose. However, upon closer examination, it becomes clear that in the majority of cases, this is far from the truth. Biofilms, especially those formed in nutrient-limiting environments, are complex, highly structured communities designed to maximize survival, reproduction and spread. They can be homogeneous, consisting of only one specific organism, or heterogeneous, comprising two or more different organisms. In nature, as in urology, both types of biofilms are commonly found and the type that will occur in any specific situation depends upon multiple factors – chiefly the properties of the surface and microorganisms present, the ability of the surrounding milieu to support and inhibit their

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growth and the relationship the organisms have with each other. As one examines these communities in more detail, it becomes evident that biofilm complexity is largely a planned behavior and not simply a matter of architecture. It would appear that every organism within a biofilm has a purpose, each contributing to the overall good of the entire population. In fact, some organisms within biofilms will undergo a type of ‘apoptosis’ or programmed cell death under certain conditions, likely to supply nutrients to the population or contribute DNA and other compounds to the matrix (Allesen-Holm et al., 2006; Whitchurch et al., 2002). Furthermore, other cells will either alter their phenotype or enter a state of dormancy (termed ‘persisters’) that renders them tolerant to lethal factors such as antibiotics (Keren et al., 2004; Lewis, 2001; Lewis, 2007; Shah et al., 2006). In this sense, a biofilm can almost be thought of as a separate entity, a multicellular organism comprising diversely differentiated cells throughout, all with a common goal (Lewis, 2001; Sauer et al., 2002). This complexity and adaptability are likely driven by the rapid environmental changes that microorganisms face on a continual basis, whether in nature or in association with human tissues and biomaterials, and offers some explanation as to why medical biofilms are virtually impossible to eradicate completely once formed. All biofilms start with an organism’s attachment to a surface (Fig. 1.1). This can be a single cell, a group of cells that grow in multiples or commonly autoaggregate, or a matrix-enclosed cluster of cells sloughed off a previous biofilm. At the time of this initial attachment, success will depend upon multiple factors, some of bacterial origin and others environmental. The

Surface approach and reversible attachment

Matrix production and irreversible attachment

Planktonic cells

Surface Conditioning film

Adhesins

Cell division, microcolony formation and biofilm initiation

Biofilm development and spread

Biofilm Matrix components Microcolony (exopolysaccharides, nucleic acids, proteins)

Water channel

1.1 Schematic of bacterial biofilm formation and development outlining the processes and stages involved.

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major bacterial determinants are surface factor expression and cell viability. Bacteria express numerous appendages, receptor ligands and polysaccharides that can become involved in adherence to a surface such as fimbriae (pili), lipopolysaccharides, exopolysaccharides, specific protein ligands and flagella (Christiansen et al., 1989; Costerton, 2007a; Jass et al., 2003; Reid et al., 1992). The possession and expression of these factors can vary widely between different types of bacteria, even down to the strain level. As the bacterium approaches the surface, one or several of these factors will contact the surface and a reversible attachment will take place, dependent upon not only those factors but also the environmental properties of the surface and surrounding fluid as well (i.e. surface hydrophobicity, energy and charge, fluid shear stress, temperature, osmolarity, etc.). This attachment allows time for additional bacterial factors to adhere and stimulates changes in gene expression within the organism that drive polysaccharide production and other events that lead to irreversible adherence (De Kievit and Iglewski, 1999; Sauer et al., 2002; Stoodley et al., 2002b; Ziebuhr et al., 1999). In the case of a biomaterial, once this irreversible attachment has taken place and a biofilm begins to form, at present it is virtually impossible to eradicate. The initial organism-to-surface attachment is further complicated during biomaterial use by the presence of biological constituents in the surrounding fluid (i.e. blood, urine), which not only provide nutrients for the microbes but also adhere to the biomaterial themselves and form what is called a ‘conditioning film’ (Fig. 1.2) (Chew et al., 2006; Choong et al., 2001; Jass

1.2 Scanning electron micrograph of a sterile urinary conditioning film formed after 24 hours on a titanium oxide surface.

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et al., 2003; Reid et al., 1992). This film negatively impacts the host in three critical ways. Firstly, it masks the intended surface properties of the device and may block the elution of an impregnated antimicrobial or anti-biofilm compound. Secondly, it provides a plethora of different molecules for the organism to attach to, which increases its ability to make and maintain initial contact with the surface. Finally, it provides a direct source of nutrients for the organism immediately where attachment and biofilm formation occur. It is mainly two aspects of the bacterium–biomaterial interaction, namely multiple bacterial attachment strategies and the constant deposition of a urinary conditioning film that have made it so difficult to produce biofilm-resistant prostheses for the urinary tract. As described previously, biofilm configuration and development are highly dependent on the local growth conditions (Costerton, 2007a; Jass et al., 2003). They can form fairly uncomplicated arrangements such as mats to highly complex spatial structures with fluid filled pores and channels (Fig. 1.3). Spatial variations can range from the bacterial scale (<10 μm microscale) to the meso-scale (50 μm–1 mm), and in the right environment they can even approach macro-scales (>1 mm). Thus, there is no one standard structure of a biofilm and a multitude of arrangements exist including mushrooms, towers, honeycombs and streamers. Furthermore, it is not uncommon to find that the same organism produces biofilms of very dissimilar structure and density under different environmental conditions (Jackson et al., 2001). In general, relatively flat and unstructured biofilms

1.3 Klebsiella pneumoniae biofilm on a ureteral stent removed from an infected patient showing water channels that permit the rapid movement of nutrients and wastes in and out of the film.

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tend to develop in nutrient-rich habitats while more highly structured, complex films form when nutrients are restricted (Costerton, 2007a). This is likely due, at least in part, to the fact that when nutrients are scarce, elaborate structures maximizing surface area and fluid flow will be required; while when ample nutrients are available in the surrounding fluid and buried within the matrix, these elaborate structures are not required.

1.3.1 Hydrodynamic effects on biofilm formation and structure The hydrodynamics of the local aqueous environment also plays a key role in determining biofilm structure. This is fundamentally due to the supply of dissolved substrates delivered by the flow, as well as the nature and magnitude of the mechanical forces imparted on the biomass by it (Nguyen et al., 2005; Stoodley et al., 1999b; Stoodley et al., 2002a). Beyenal and Lewandowski (2002) proposed that biofilms arrange their internal structure according to the velocity of the flow they are grown in. They discovered that biofilms grown in high velocity gradient flows would adapt their internal structure to bolster their resistance to elevated shear stress. The strength increase comes about with an increase in biofilm density which slows down the internal mass transfer rate, thus reducing the nutrient transfer to deeper layers. Similarly, biofilms grown at low flow velocities evolve higher effective diffusivity but are not as dense and subsequently exhibit a weaker resistance to shear stress. In addition to biofilm formation, hydrodynamic forces can influence biofilm migration. In a study of Pseudomonas aeruginosa biofilms, Purevdorj et al. (2002) depicted a unique mechanism of biofilm locomotion developed under hydrodynamic stress. They demonstrated that biofilms could move along solid surfaces while remaining attached to these surfaces. This was unlike the more commonly known propagation process via fluid-borne detached planktonic (free floating) cells; this surface-associated mechanism allowed the spread of the whole biofilm structure. With diffusion of the entire structure comes the preserved resistance to various antibiotics and disinfectants, since the intermediate step of planktonic release is not required (Stewart and Costerton, 2001). Such a creeping biofilm flow may be an important consideration for infection in catheters and stents. Formation of a biofilm often represents a flow impediment. The interaction of flow and biofilm has most commonly been modeled with the biofilm simulated as a solid object around which the flow must navigate. Fluids indeed flow around biofilm structures but flow also exists through the extracellular polymer substance (EPS) matrix and microchannels that can form within the film (Fig. 1.3) (Nguyen et al., 2005; Stoodley et al.,

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1994). This multi-scale environment can make for subtle but potentially significant complications in models that attempt to simulate the flow around and through biofilms. Microscopic flow effects through biofilms can affect the local mesoscopic flow field (Nguyen et al., 2005). The challenge in fully quantifying the effects of flow in, at, and around biofilms is largely a result of the complex mechanical material properties that a biofilm exhibits. The hydrodynamic impact of biofilm accumulation is well documented, particularly in the infrastructure engineering realm. In general, hydrodynamic resistance is a function of the frictional resistance of the fluid boundary (surface drag), and the losses due to a difference between upstream and downstream pressures around an obstacle – caused by boundary layer separation (Debler, 1990). Thus, in the context of biofilms we may encounter increased skin friction and form friction from biofilm structural geometry (ripples, dunes, streamers) (Stoodley et al., 1999a; Towler et al., 2007). Even the simplest flat biofilm geometry can lead to urological implications, specifically with regard to flow through implants. The resistance to idealized flow through a circular conduit is proportional to the fifth power of the conduit diameter. That is to say, hydraulic losses are very sensitive to reductions in diameter. Reductions in device diameter owing to biofilm longitudinal and circumferential propagation may mean reduced urinary flow rates should the necessary compensatory increase in pressure gradient to combat losses across the device not be available. In considering the local environmental conditions of the urinary tract, the above findings would suggest that relatively flat films possessing high effective diffusion rates would tend to develop on urinary prostheses, as ample nutrient levels and low flow velocities generally exist immediately around them. Studies investigating ureteral stents inserted for up to 128 days support this, finding that biofilms formed in the absence of encrustation did not induce blockage of any of the devices, even in patients whose biofilms developed over several months (Reid et al., 1992). Furthermore, those devices examined using scanning electron microscopy revealed relatively flat biofilms with depths in the micrometer range. Thus, in the majority of cases, it would appear that the major clinical difficulties associated with urinary biofilms are not due to device failure but more to biofilm-associated infections and their resilience. However, this would not be the case for biofilms associated with urease-producing organisms, since the rapid alkalinization of the urine they induce would lead to rapid encrustation and subsequent blockage of the device due to the formation of more crystalline biofilms with higher densities (Morris et al., 1999; Stickler et al., 1998). These organisms, particularly Proteus mirabilis, are covered in more detail in Chapter 7, Proteus mirabilis biofilm formation and catheter design, by David Stickler, an expert in the field.

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1.3.2 Communication within biofilms The success of microorganism survival in general is largely owed to their ability to sense even minute environmental changes and make rapid adjustments. In order to achieve this, microorganisms have developed a multitude of ligand–receptor and sensor–response regulator systems for assessing their local conditions and eliciting a response (Khorchid and Ikura, 2006; Mascher et al., 2006). These triggers can work alone or in combination and may result in the organism altering a number of biological processes including growth rate, movement, membrane permeability, efflux rates, exopolysaccharide production and cell lysis. With this in mind it is understandable that microorganisms, especially those within biofilms, would constantly assess their own population dynamics and adjust their phenotypes accordingly. The best characterized example of cell–cell communication is that of quorum sensing in the marine bacterium Vibrio fischeri, which resides in a symbiotic relationship within the light organs of several marine creatures (Dunlap, 1999). When a critical number of organisms (quorum) is achieved within the organ, a factor secreted by the bacterium (inducer) builds up to a level where it binds to a sufficient number of surface receptors to trigger the expression of a bioluminescent protein important to the fish for avoiding predators at night. These inducer–receptor systems and/or the like have since been found in virtually all microorganisms tested and have been shown to regulate a multitude of biological processes including virulence and biofilm formation, with Pseudomonas aeruginosa the most studied (Christensen et al., 2007; Erickson et al., 2002; Murray et al., 2007; Rumbaugh et al., 2000). Typical quorum-sensing compounds for Gram-negative and -positive bacteria are acylated homoserine lactones (Erickson et al., 2002; Hentzer et al., 2002) and processed oligo-peptides (Abraham, 2006), respectively. Multiple studies have further argued that quorum-sensing systems may be used specifically by biofilm organisms to mobilize to more amenable locations within the film (Hansen et al., 2000), modify the local environment (Allison and Gilbert, 1995), and attract and detach organisms to and from the biofilm, respectively (Stoodley et al., 2001a; Stoodley et al., 2001b). Another potential means by which cell–cell signaling occurs within biofilms is through electrical pulses sent along microbially produced ‘nanowires’. The production and utility of such structures was first observed in biofilms of Schewanella oneidensis (Gorby et al., 2006), a versatile, environmental organism infrequently isolated from human opportunistic infections. These nanowires have since been demonstrated in Geobacter sulfurreducens (Reguera et al., 2006) as well, a metal-reducing bacterium with multiple bioremediation applications. From this it can be hypothesized that these structures are not limited to these organisms and may instead be a fairly common characteristic of biofilms. This could allow for the rapid transfer

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of both energy and information throughout the biofilm, potentially resulting in global or partitioned phenotypic changes.

1.4

Biofilms in general medicine

The human body is covered with bacteria, inside and out, such that our own cells are roughly outnumbered 10 to 1. It is largely due to this fact that the human body has evolved multiple strategies to deal with these potential invaders, from bodily mechanics to the development of a complex immune system, as well as the procurement of commensal microorganisms to inhabit our skin and oral, respiratory, gastro-intestinal and urogenital tracts. Despite these tactics, biofilms remain an extremely difficult mechanism for humans to cope with in terms of both pathogenic and opportunistically pathogenic microorganisms. The CDC (Centers for Disease Control and Prevention) reports that approximately 65% of all infections in developed countries are caused by biofilms, making apparent the fact that biofilm formation plays a key role in many clinical and sub-clinical infections (Hall-Stoodley et al., 2004). Biofilms have been identified in virtually every system in the human body, and prosthetic devices – including artificial joints, urinary catheters and heart valves – are prone to the formation of biofilms once exposed to pathogens and present a particular dilemma for clinicians (Gristina, 1987). In this section we briefly review some occurrences of biofilm formation in clinical infection in general. Biofilms are primarily associated with bacterial infections, either in association with human infections or adherent to the surfaces of medical devices. That said, other pathogens have been known to form biofilms including yeasts and other fungi (Saarela et al., 2004). For instance, Candida spp. biofilms form in a similar fashion to those of bacterial species, and as with bacterial biofilms, fungal biofilms provide microbes with enhanced resistance to medical treatment with antifungal agents (Jabra-Rizk et al., 2004). The most commonly encountered biofilms in the human body reside in the oral cavity in the form of dental plaques (Marsh, 2003). In addition, chronic bacterial infections of the head and neck are commonly associated with biofilm formation. Such infections include chronic otitis media and chronic tonsillitis as well as prosthetic infections associated with endotracheal tubes and voice prostheses (Post et al., 2004). Chronic otitis media was long thought to be an inflammatory condition secondary to the fact that only one-third of infections had positive cultures. Evidence now suggests that bacterial biofilms form on the inner ear and occasionally release planktonic bacteria resulting in sporadic positive cultures (Post, 2001). This is supported by the finding that many biofilm-associated organisms commonly resist cultivation using standard plating techniques and media (Veeh et al., 2003). Foreign materials such as endotracheal and tracheostomy tubes

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present surfaces that are prone to biofilm formation. Natural host defenses such as coughing are rendered ineffective by the presence of the device and biofilm formation provides antibiotic resistance leading to complications such as ventilator acquired pneumonia. Wound infections may be separated into acute and chronic infections. Acute wound infections are encountered in situations such as surgical wounds where pathogens are introduced into the tissues at the time of surgery. Despite preventative measures such as preoperative skin preparation with antimicrobial solutions, vigorous hand washing and sterile surgical technique it is estimated that roughly 3% of surgical wounds become infected (Horan et al., 1993). It is likely that virtually all surgical incisions are contaminated with bacteria, primarily with Staphylococcus spp. from the skin; however relatively few of these wounds become clinically infected. Factors such as the site of surgery (i.e. skin, bowel), skill of the surgical team, host factors and the virulence of the bacteria determine whether wound colonization results in infection (Fry and Fry, 2007). Chronic wound infections such as those associated with diabetic foot ulcers and venous leg ulcers are a significant cause of morbidity and mortality around the world. Infection of these wounds is well known to impede the process of wound healing and adequate antimicrobial therapy is essential to prevent complications such as sepsis or limb loss. It has been shown that many chronic infections result from bacterial biofilm formation, a process that likely contributes to the difficulty in their treatment (Costerton et al., 1999; Donlan and Costerton, 2002; James et al., 2008; Percival and Bowler, 2004). In general, evidence suggests that while biofilms are prevalent in chronic wounds they are relatively rare in acute wounds (James et al., 2008). Burn injuries present a different sort of acute wound that may result in life-threatening infection. Death from systemic sepsis is the most common cause of mortality among burn patients (Mayhall, 2003). Thermal injuries are easily colonized by organisms such as Staphylococcus spp. and Pseudomonas aeruginosa. Strategies to prevent and/or treat infected burn wounds include systemic and topical antimicrobial treatments and dressings. There is evidence to support the role of biofilm formation in burn wounds enhancing bacterial resistance to treatment. Efforts have been made to develop topical therapies to both reduce biofilm formation and provide bactericidal action for the treatment of burn infections (Martineau and Dosch, 2007). Patients with valvular heart disease, both native and prosthetic, have also been the focus of significant biofilm research. Native valve endocarditis results from interaction between the endothelial lining and circulating bacteria, primarily streptococci, although fungi have been identified as well. Additionally, infection of prosthetic heart valves (prosthetic valve endocarditis (PVE)) may affect up to 4% of patients with these devices (Douglas and Cobbs, 1992). Colonization of bacteria is most likely to occur at the

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sights of implantation of these valves where the tissue has been injured. PVE may be divided into early and late, with early infections being primarily caused by coagulase-negative staphylococci thought to result from contamination. Late PVE (>12 months) may also include S. aureus, enterococci, and fungi among pathogens (Donlan and Costerton, 2002). As with endotracheal tubes and prosthetic heart valves, virtually every type of prosthetic used in the human body is prone to biofilm formation and infection. Strategies to prevent or eliminate biofilms will be discussed in greater detail later in this chapter. As discussed, the most commonly employed strategy to date, once prosthetic infection has been identified, is the complete removal and replacement of the prosthetic device. The disadvantages of this approach include the greater risk to the patient with increasing surgery, the risk of persistent infection and the greater technical difficulty of replacing the prosthetic in a previous surgical site. As we turn the discussion to the importance of biofilms in genitourinary infections, it will become increasingly important to distinguish between biofilms associated with prosthetic versus non-prosthetic infections.

1.5

Biofilms in urology

1.5.1 Hydrodynamics of the upper and lower urinary tract In the upper urinary tract, urine flows largely peristaltically from the kidney to the bladder. Although modern research on this process has been active since Boyarski (Boyarsky et al., 1971), many details of urine transport remain unresolved (Vahidi and Fatouraee, 2007). This may be a result of the complexity of the transport process, which is not completely peristaltic. A second component of the flow is driven purely by a pressure gradient developed between the renal pelves and the bladder. This second flow component can be problematic should pressures in the bladder exceed those at the renal pelves. Furthermore, computational analysis has demonstrated that the normal initiation of peristalsis can itself momentarily reverse the flow in segments of the ureter nearest its upstream entrance (Vahidi and Fatouraee, 2007). These potential reflux conditions can lead to a reverse flow of urine, along with potential bacteria and toxins, from the bladder to renal pelves and on to the kidneys (Bykova and Regirer, 2005). Since the placement of a ureteral stent relaxes the annular muscles at the stent–ureter interface, subsequently disabling the peristaltic urine transport, it can further encourage this backward flow. Under these conditions urine flow, through and around the stent, becomes completely driven by the pressure differential that exists between the pelves and the bladder. This reduced and at times retrograde flow enhances the ability of biofilms to migrate along the device toward the kidneys, potentially leading to pyelonephritis

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(Choisy et al., 1998; Choong and Whitfield, 2000; Cummings et al., 2004). Importantly, it has been shown in mathematical modeling that perforating holes common to most stents act to reduce this potential for reflux (Cummings et al., 2004). Recent numerical modeling of a stented ureter revealed that these perforations or pass-through holes are generally inactive during flow, only to become stimulated once there is a fouling or blockage (Tong et al., 2007). In general, urine flow from the bladder via the urethra has received little theoretical attention (Bykova and Regirer, 2005). Only relatively recently (Lecamwasam et al., 1999), experiments on surgically separate canine urethras revealed that urethral flow was most likely laminar. It has been speculated that the flow may briefly or sporadically transition to a non-laminar state in humans. Common flow rates in the urethra vary between 9 and 21 ml/s in adult males and between 10 and 18 ml/s in females. With the implementation of a catheter the characteristics of lower urinary tract flow change substantially. At the downstream end, the catheter expels the urine outside of the body; at the upstream end, the catheter extends into the bladder. By way of extension into the bladder the catheter effectively holds the internal sphincter open so as to create a constant drip flow through the catheter to a collection device. This sort of low energy flow has negative consequences for the patient as the higher shear rates associated with normal urination are no longer present. High flow velocities in normal urinary flow are a result of the driving pressure built up in the bladder by a closed internal sphincter. The higher shear rates connected with these flows are often successful at flushing out many of the potentially detrimental resident bacteria. A further catheter complication stems from the balloons attached to maintain catheter position. Often, the balloon and catheter orientation can lead to the formation of a urine sump at the bottom of the bladder leading to an accumulation of bacteria and deposits. In an attempt to regain some of the advantages of the healthy urine flushing regimen, some research has focused on the development of an electronic catheter valve. The valve can be opened on a programmable time release basis or manually by the patient through a proximity activator worn as a ring on a finger. Clinical trials of this device are due to finish around 2009 (EvansPughe, 2005). Clearly, this is an area where substantial gains can be made for patient comfort with innovative solutions that tap microbiological, urological and engineering expertise.

1.5.2 Urinary infections associated with biofilms Infection of the genitourinary tract has long been a source of frustration for patients and physicians. Chronic bothersome infections such as recurrent cystitis and chronic bacterial prostatitis often resist treatment despite

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multi-modality therapy. It is certainly reasonable to assume that the formation of biofilms in the genitourinary tract is, in part, responsible for the chronicity of these infections as well as their resistance to treatment. In this section we examine some of the basic factors involved with genitourinary infection in general, as well as several specific examples where biofilms likely play a significant role in the infectious process. The normal human urinary tract contains multiple factors to protect against infection including the antegrade flow of sterile urine, the presence of commensal bacteria and a smooth epithelial lining (urothelium). A number of additional host strategies exist that will be discussed in greater detail further into the chapter (Section 1.7.1). The failure of any or all of these systems through aging or pathologic processes may render the host more open to pathogenic bacterial colonization and infection. For example, the distal 1 cm of the female urethra contains a biofilm composed predominantly of lactobacilli similar to those present in the vagina (Costerton, 2007b). In this setting the biofilm is actually beneficial and functions to prevent the passage of pathogenic bacteria from the environment to the bladder. The distal portion of the male urethra contains a similar biofilm barrier and is significantly longer than the female counterpart making colonization even less likely (Costerton, 2007b). Any process that compromises this initial biofilm defense may eventually lead to urinary tract infection. Such processes include the insertion of foreign bodies (i.e. urethral catheters), colonization with pathogenic bacteria and the use of broad-spectrum antimicrobials that destroy the commensal bacteria. As we turn to specific examples of genitourinary infection involving biofilms, we will distinguish between those infections involving prosthetics (i.e. penile prostheses, urinary catheters, ureteral stents) and those that do not (i.e. chronic bacterial prostatitis, staghorn calculi). Chronic prostatitis (CP) and male chronic pelvic pain syndrome (CPPS) present a challenging management dilemma for clinicians. Many patients suffer with ongoing symptoms such as urinary frequency and urgency as well as pelvic and penile pain without ever conclusively discerning the cause. Patients with such symptoms are often treated with antimicrobial agents even in the absence of positive cultures in an effort to eradicate sub-clinical infection. The leading theories behind the evolution of bacterial prostatitis are the retrograde spread of bacteria from the environment through the urethra and the reflux of infected urine through prostatic ducts. Acute bacterial prostatitis may be well treated with antibiotics while pathogens have not formed colonies and biofilms. Common pathogens include Escherichia coli, Klebsiella spp., enterobacteria, Pseudomonas aeruginosa, Staphylococcus spp. and Proteus spp., among others (Donlan and Costerton, 2002). Studies in animal models of prostatitis suggest the formation of glycocalyx-encased microcolonies of bacteria adherent to

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prostatic mucosa layers (Nickel et al., 1990). Nickel and Costerton (1993) examined prostatic biopsies in men with a history of chronic bacterial prostatitis. Bacteria were found attached to prostatic ductal walls, particularly P. aeruginosa. Others studies have clearly demonstrated the presence of biofilms in men who failed treatment for chronic bacterial prostatitis (Domingue and Hellstrom, 1998). It is a logical conclusion that biofilm formation is, at least in part, responsible for the failure of treatment in many men with chronic bacterial prostatitis. In the future, strategies for the prevention or disruption of these biofilms may significantly enhance our ability to manage this challenging disease. Staghorn renal calculi are stones that fill much of the collecting system of the kidney and may result in serious complications such as urinary obstruction, renal failure or sepsis. These stones are typically associated with urea-splitting bacteria (primarily Proteus spp.) and, if left untreated, may result in renal loss or even death. Proteus vulgaris produces biofilms and struvite crystals that may lead to petrification of the urinary collecting system (Costerton, 2007b). Patients with staghorn calculi frequently have recurrent urinary tract infections that respond initially to antibiotics but almost certainly recur so long as the stone remains untreated. Successful treatment of this condition requires complete removal of all visible stone, usually through percutaneous renal surgery followed by antibiotic therapy to sterilize the urine and prevent recurrent infection and biofilm formation. In North America alone more than 100 million urinary tract devices (urethral catheters, ureteral stents, penile prostheses) are inserted each year, translating into millions of device-associated infections and billions of dollars in additional health care expenditure annually (Jacobsen et al., 2008). One of the major problems associated with these foreign bodies is that they present novel, non-host surfaces on which bacteria can colonize and form biofilms. Urethral catheters are commonly used during surgery to monitor urinary output and do not typically result in clinical infection as there is a sufficiently short indwelling time and bacteria are unlikely to form biofilms. Conversely, many patients require long-term urethral catheterization and may suffer from chronic recurrent bladder infections. It has been estimated that the risk of bacteriuria increases by roughly 10% for each day that a urethral catheter is in place and that virtually all catheterized patients will develop urinary tract infection when the catheter remains in place beyond 28 days (Stickler, 1996). Not only does the colonization of urinary catheters lead to urinary tract infection but it may also affect the functioning of the catheter through obstruction of the lumen with encrustation (Liedl, 2001; Stickler et al., 1993; Stickler, 1996). Bacteria spread in a retrograde fashion and tend to progress faster through the lumen of the catheter than on the external surface (Nickel et al., 1985). Ureteral catheters (i.e.

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ureteral stents) are likewise subject to biofilm development, infection and encrustation. These devices may be colonized despite sterile urine cultures, again drawing attention to the lack of success in cultivating biofilm organisms (Lifshitz et al., 1999). The major bacterial species involved in catheter and stent infections are Escherichia coli, Proteus mirabilis, Klebsiella pneumoniae, Pseudomonas aeruginosa, Staphylococcus epidermidis, Enterococcus faecalis and Staphylococcus aureus. Ureteral stent biofilms involving several of these pathogens are shown in Fig. 1.4. In addition, there are countless opportunistic pathogens that can form biofilms on these devices and cause infection depending upon the patient’s health and immune status, as well as microbial exposure. Management of patients with long-term catheterization and stenting remains a challenge. In most situations, only patients at risk for sepsis or those with symptomatic urinary tract infections should be actively treated while those with isolated bacteriuria may be observed. Routine device changes may help to reduce bacterial load and allow for more effective antibiotic therapy when required, since mature biofilms may be more difficult to treat than those less established (Anwar et al., 1989). Genitourinary prosthetic devices such as artificial urinary sphincters and penile prostheses are convenient targets for bacterial seeding and biofilm infection (Silverstein and Donatucci, 2003). This is typically far more devastating to the patient than catheter infection due to the possible need for device removal and replacement. Not only are these surgeries difficult, but reoperations for prosthetic devices are associated with a much higher risk of infection than a primary procedure. It is accepted that most prosthetic infections result from bacterial seeding during surgery (Silverstein and Donatucci, 2003). These bacteria may cause early clinical infections or may form biofilms on the device surface and take years before causing a clinically apparent infection. Staphylococcus epidermidis is commonly isolated from infected penile prostheses and may be present in up to 80% of these infections (Blum, 1989). Interestingly, it is likely that the majority of penile prosthetics are seeded and colonized although most will never manifest with clinical infection. Silverstein et al. (2006) examined the devices of ten patients undergoing device revision or replacement of clinically noninfected penile prostheses. Bacterial biofilm formation was identified in eight of ten prostheses with multiple associated organisms suggesting that the number of prostheses colonized with bacteria is far greater than was previously believed. The treatment of infected prostheses has traditionally involved device removal with a course of antibiotic therapy and delayed reinsertion of a new prosthetic (Silverstein and Donatucci, 2003). Infection is associated with pronounced inflammation and corporal scarring making insertion of a new device extremely difficult (Wilson and Delk, 1995). In more recent

Introduction to biofilms in urology (a)

(b)

(c)

(d)

(e)

(f)

19

(g)

1.4 Panels show biofilms formed as a result of infection with (a) Candida albicans, (b) Enterococcus faecalis, (c) Escherichia coli, (e) Klebsiella pneumoniae and (f) Staphylococcus aureus; panel (d) shows encrustation caused by Proteus mirabilis and panel (g) is a crosssection of a ureteral stent infected with E. coli and P. mirabilis.

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years, it has become far more common to attempt salvage of infected penile prostheses with removal of the infected device and replacement of a new device during the same procedure with adjuvant antibiotic treatment (Brant et al., 1996). From the standpoint of biofilms this makes sense as device removal should eliminate the vast majority of established biofilms and remaining bacteria are likely to be in the planktonic state where they will be more susceptible to antimicrobials.

1.5.3 The female urogenital tract The female urogenital tract is an interesting microbial environment, in that it consists of areas commonly inhabited by microorganisms (vagina and distal urethra) and others regarded as being sterile (upper genital and urinary tracts). Thus, while generally 107–108 colony forming units (CFU)/ gram of fluid are present within the vagina, the bladder and uterus are, for the most part, devoid of organisms. However, since organisms residing within the vagina have access to both the upper genital and urinary tracts, the vaginal microbiota can greatly influence the development of infections throughout the entire tract. In addition, factors such as sexual relations and exposure to spermicides, pregnancy, hormone fluctuations, menstruation, hygienic procedures, antibiotic usage and constant exposure to a vast array of environmental and fecal-derived organisms make it a highly dynamic region for both the host and microbes (Reid et al., 1990). This relative instability suggests that a niche will occur at times in which potential pathogens can attach, grow, form biofilms and cause infection. Indeed, aside from vaginal infections, studies have demonstrated a significant correlation between bacterial colonization of the vagina and periurethral area and the development of urinary tract infections (Stamey et al., 1971; Stapleton et al., 2002). In a normal, healthy vagina the pH is generally below 4.5, which largely restricts the organisms that can survive there. Understandably, acidophilic species of bacteria, such as lactobacilli, typically predominate. Indeed, members of the Lactobacillus genus are the most commonly isolated organisms from the vaginal microbiota of healthy women, and as such have been hypothesized to play a major role in maintaining this healthy status. When the numbers of vaginal lactobacilli drop, typically due to one of the events listed above, multiple types of infection can develop. These include yeast vaginitis, bacterial vaginosis (BV), aerobic vaginitis (AV) and urinary tract infections. Collectively, these infections cover the fungi as well as most non-sexually transmitted bacterial species known to inhabit the urogenital tract. While BV is predominantly associated with anaerobes – such as Gardnerella vaginalis, Mobiluncus spp., Bacteroides spp., Peptostreptococcus spp., Prevotella bivia, Mycoplasma hominis, Atopobium vaginae and Ureaplasma

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urelyticum (Hale et al., 2006; O’Brien, 2005) – AV is generally caused by multiple species of aerobic bacteria including Escherichia coli, Staphylococcus aureus and members of the Group B streptococci (Romanik et al., 2007). Of these, only biofilms of G. vaginalis have been extensively studied within the vagina and been shown to regularly occur (Swidsinski et al., 2005). However, the frequent recovery of bacteria-laden epithelial cells from vaginal swabs as well as the ability of many of these organisms to remain in the tract for extended periods of time strongly supports their ability to form biofilms as well (Hillier, 1993; Saidi et al., 1994; Sobel, 1990). These biofilms have the potential to be powerful nidi for infection, especially in the case of urethral catheters, as aseptic technique during device insertion has no bearing on potential colonization. Ultimately, to maintain overall urogenital tract health in women, the goal is to simultaneously support the growth of commensal Lactobacillus biofilms within the vagina while inhibiting those associated with pathogens.

1.5.4 Intracellular biofilm communities (IBCs) caused by uropathogenic escherichia coli (UPEC) It has been known for over 25 years that certain strains of E. coli could enter urinary epithelial cells (Fukushi et al., 1979), and it is now well recognized that this mode of attack is quite common during infection with UPEC. The host’s response to these organisms is exfoliation of the bladder epithelial lining to remove these bacteria-laden cells while preparing both innate and acquired immune response modalities to deal with the remaining organisms still present. However, it was not until Scott Hultgren’s group discovered the ability of UPEC cells to migrate through the superficial umbrella cells and into underlying layers of the bladder, where they multiplied and formed bacterial communities, that the idea of intracellular UPEC biofilms was formed (Anderson et al., 2004). While all of the early work by Hultgren’s group was performed using a murine model of UPEC cystitis (Anderson et al., 2004; Justice et al., 2004), their most recent studies have demonstrated these same IBCs in the exfoliated bladder cells of women with active cystitis or a history of its recurrence (Garofalo et al., 2007; Rosen et al., 2007). This finding offers the most likely explanation for observed UPEC resistance to host immune factors and antibiotics, in addition to their propensity for inducing chronic infections in the absence of immune defects or urinary prosthetics (Mysorekar and Hultgren, 2006). Furthermore, these findings offer potential insight to other conditions and chronic infections which may use similar mechanisms. One such urinary condition is interstitial cystitis, which has been suggested to be caused by chronic infection with undetected organisms such as Enterococcus spp., despite little solid

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evidence (Elgavish et al., 1997). Future research is required to determine whether any connections actually exist.

1.6

Biofilm shedding and migration: infection spread and recurrence

Detachment and dispersal of biofilm cells into the environment is the least known aspect of biofilm formation but is critical to bacterial spread and thus, overall population survival (Costerton, 2007a; Costerton et al., 1987; Stoodley et al., 2002b). A natural example of this is a bacterial biofilm located on a rock within a stream constantly shedding off cells which begin new biofilms on rocks further downstream. In terms of a surgical wound or biomaterial use, biofilm shedding can lead to biofilm and infection spread, sepsis, the reappearance of infection following antimicrobial treatment (often leading to chronic infection) and the growth of the original biofilm. For example, bacterial cells released from a ureteral stent biofilm have the potential to: (a) travel down to the bladder and instigate or regenerate a previously cleared cystitis; (b) travel up the device itself (i.e. Proteus mirabilis swarming, biofilm migration) (Purevdorj et al., 2002; Stickler, 1999) or, via urine reflux induced by the device, travel to the kidney, causing pyelonephritis (Choisy et al., 1998; Choong and Whitfield, 2000; Cummings et al., 2004); (c) adhere again to the device and/or existing biofilm increasing its surface area and density; (d) adhere to a potential calculus within the tract, creating a further nidus of infection and potentially inhibiting stone treatment options; (e) promote the formation of calculi if the organism expresses urease (Choong et al., 2001; Stickler et al., 1993); and (f) enter the underlying layers of the urinary tract epithelia and even the bloodstream if local damage has been induced by the device or infection, or the organism possesses the ability to do so on its own (Garofalo et al., 2007; Justice et al., 2004; Rosen et al., 2007; Wright et al., 2007). All of these scenarios enhance the ability of the organism to remain in the tract following device removal and/or replacement, and can greatly increase the severity of infection. Cadieux et al. (2008) compared a standard ureteral stent placed in conjunction with pre- and post-operative antibiotics to a stent impregnated with the antimicrobial triclosan in chronically stented patients. They observed that every stent placed in a patient with a positive urine culture during instillation had adherent, viable biofilms upon removal at 3 months, despite the fact that they all went through multiple courses of antimicrobial therapy during the indwelling period. Furthermore, biofilm organisms from every one of these devices matched the original urinary isolate as well as any organisms recovered from their previous stent(s). This strongly suggests that ureteral stents placed within an infected urinary

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environment will become infected, regardless of the device type or antimicrobial therapy regimen used, and highlights the ability of biofilms to induce chronic infections. Ultimately, from a clinical standpoint, bacterial detachment is likely the most negative aspect of a biofilm as it provides organisms that not only enhance, spread and prolong infections but can also disseminate and survive in the host in numerous ways until a new device is inserted.

1.7

Resistance to host factors and antibiotics

1.7.1 Host factors As briefly described earlier in the chapter, the urinary tract is under continual assault from microorganisms due to its direct connection to the outside world, constant production of nutrient-rich urine and close proximity to the ever-bacteria-releasing intestinal tract. It is therefore understandable that the vast majority of human urinary tract infections are caused by fecally derived organisms, following their migration along the perineal skin, transfer by clothing, poor hygienic practices, sexual contact or migration from the vagina in the case of women. In response, the urinary tract has developed numerous strategies to combat these invaders. Micturition (voiding) promotes high shear stress, helping to flush out microorganisms from the bladder and urethra that are planktonic or loosely adherent to the epithelia. In addition, at each voiding the bladder is drained completely, which helps in expelling as many of the organisms present as possible. Exfoliation of the superficial umbrella cells of the bladder epithelia removes strongly adhered or intracellular organisms on or in the outer layer of the bladder wall. The expression of Tamm–Horsfall protein helps block the Type 1 fimbriae FimH receptors on Escherichia coli to prevent their attachment to mannosylated receptors on tissues and promotes bacterial clearance upon voiding through aggregation. Immunoglobulin A (IgA) antibodies and pro-inflammatory cytokines and chemokines opsonize organisms and recruit and stimulate immune cells such as neutrophils and macrophages to the tract, respectively, to actively destroy pathogens. Cationic peptides termed ‘defensins’ are produced by neutrophils as well as kidney epithelial cells, where they are secreted into the urine at concentrations of approximately 10–100 μg/ml (Ganz, 2001). These peptides form multimeric pores on the surfaces of fungi as well as Gram-positive and -negative bacteria, resulting in efflux of cellular contents and intracellular dysregulation that lead to bacterial killing (Lehmann et al., 2002). Finally, even if a symptomatic infection does develop that is not resolved by host defenses, numerous antibiotics are available that can help the immune system eliminate it with few or no serious side effects.

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Thus, it is clear that the host urinary tract has evolved a sophisticated series of mechanisms through which to prevent and eradicate potential infections. However, following biomaterial insertion in to the tract, many of these mechanisms become compromised, especially if microorganisms achieve access to the device and develop a biofilm on its surface. Firstly, the presence of the device itself can impair complete voiding and induce inflammation and local tissue damage, physically disrupting normal host microbial barriers. In addition, strong adherence of organisms to the device prevents their eradication during voiding and exfoliation. The biofilm matrix renders buried bacteria less accessible by antibodies, neutrophils and macrophages and can prevent or slow the movement of antimicrobial peptides and hydrogen peroxide through the biofilm (Lewis, 2007; Vuong et al., 2004). This prevention or delayed movement of antibacterial factors through the biofilm can give organisms an opportunity to engage their own defense strategies and can also result in only sub-lethal concentrations reaching their target. Finally, metabolically slowed and/or dormant persister cells within the biofilm are largely tolerant of antimicrobial substances released by the host, even when at levels that are lethal to planktonic cells (Cernohorska and Votava, 2002; Fedtke et al., 2004; Leid et al., 2002; Lewis et al., 2007).

1.7.2 Antibiotics The fact that biofilms are inherently resistant to antibiotics has been well documented and is the greatest obstacle in treating chronic infections, especially those associated with prostheses. As a consequence, infections associated with biofilms are rarely resolved using antibiotic therapy alone, even in patients with fully functioning immune defenses. Studies have shown that some organisms within biofilms can survive antibiotic concentrations >1000fold higher than those required to kill their planktonic counterparts (Anderl et al., 2000; Stewart and Costerton, 2001). This resistance is multi-factorial, dependent upon the organism challenged and agent utilized, the structure and makeup of the biofilm matrix, local environmental factors (temperature, hydrodynamic forces, nutrient levels, etc.) and the apparently constant generation of persister cells. It was originally hypothesized that the EPS matrix of biofilms was the primary factor involved in their resistance to antimicrobials, strictly owing to the inability of the antimicrobial agent to reach cells buried within the biofilm. Although this is true to an extent, several studies have also demonstrated rapid transport of certain agents through biofilms, via water channels and diffusion through the EPS matrix itself. To clarify, while the matrix can inhibit the transfer of particular compounds including antibiotics through it via chelation, destruction and direct blockage, it is largely agent

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and biofilm dependent (Brady et al., 2006; Dasgupta et al., 1997; Jefferson et al., 2005; Nichols et al., 1988; Ramphal et al., 1988). Ultimately, these impediments may result in only sub-lethal concentrations of the antibiotic reaching various cells within the biofilm population or delay the buildup of the compound to cidal levels, allowing resistance mechanisms to be activated or acquired. Examples of this include biofilms formed by strains expressing beta-lactamase (Anderl et al., 2000) (destroys most penicillin class antibiotics), and the chelation of aminoglycoside antibiotics by negatively charged particles within the biofilm matrix (Nichols et al., 1988). Another proposed mechanism to explain biofilm-associated antibiotic resistance is the development of various microenvironments within certain pockets of the biofilm that may be anaerobic, have altered pH and/or vary highly in solute and waste product levels. These changes may result in antibiotic destruction in that region or induce phenotypic changes in some cells that result in a slow-dividing state or reduced membrane permeability. This is supported by studies showing reduced antibiotic susceptibility in membrane porin mutants (Otto and Hermansson, 2004), loss of activity of some antibiotics under anaerobic conditions (Tack and Sabath, 1985) and the fact that several classes of antibiotics, such as the beta-lactams (which inhibit peptidoglycan synthesis), are only effective in rapidly dividing cells. A third mechanism of resistance is the production of the aforementioned persister cells, a sub-population of metabolically dormant cells within biofilms that are akin to bacterial spores (Keren et al., 2004; Lewis, 2007; Shah et al., 2006). These cells are highly tolerant to antimicrobials and appear to be produced maximally during stationary phase growth, suspected to be the phase of growth in which most biofilm cells reside. They do not apparently need to possess any particular resistance mechanisms to survive and can resume growth once the concentration of antibiotic drops to sub-lethal concentrations (Lewis, 2007). Although planktonic persisters can be mopped up by immune cells following antimicrobial therapy that kills the majority of the remaining population, cells buried within the biofilm matrix are protected from immune clearance. This explains how persisters can be formed during both planktonic and biofilm infections but only become a clinical problem with respect to biofilms.

1.8

Current and future biofilm prevention and treatment strategies

The prevention and treatment of biofilms in urology, especially those on prostheses, has involved multiple research and clinically based approaches. Most strategies have concerned either modifications to the properties of

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the biomaterial itself (Chandra et al., 2005; Legeay et al., 2006; Petrini et al., 2006; Poncin-Epaillard and Legeay, 2003; Rupp et al., 2002; Schierholz et al., 2001) or the development and application of surface coatings (Hou et al., 2007; Vigeant et al., 2005), in attempts to prevent bacterial attachment by rendering the biomaterial surface less attractive to microorganisms. This line of investigation has evolved further to include the use of adherent, impregnated and oral antimicrobials to kill or inhibit the growth of organisms that are ultimately successful in gaining the surface (DiTizio et al., 1998; Nablo et al., 2005; Norris et al., 2005; Petrini et al., 2006; Reid et al., 2001; Yoshinari et al., 2006). Although no single approach has proven entirely successful, several have demonstrated some clinical benefit in reducing infections (Davenport and Keeley, 2005; Reid et al., 2001; Slawson et al., 1992), and many have shown superior efficacy during in vitro and/or animal studies. As mentioned earlier in the chapter, the contributing factors that preclude complete biofilm prevention and/or clearance are multiple bacterial attachment strategies, the rapid and consistent development of a host-based conditioning film on all implanted devices, organism protection by the biofilm matrix, tolerance and resistance to antimicrobial factors and the impaired immunity of many of the patients. Although no biofilmresistant biomaterial or device currently exists in urology, we will briefly review several major avenues of research aimed at developing one.

1.8.1 Antimicrobials Silver The most studied and arguably successful surface-based strategy to date for preventing device-related infections in urology is that of silver-coated urethral catheters, which have been in use for over 20 years. The broadspectrum antimicrobial effects of silver have been known since ancient times, and detailed studies have revealed that it acts by blocking multiple enzymatic processes via interaction with electron donors (especially sulfhydryl groups) (Slawson et al., 1992) and through membrane destabilization (Dibrov et al., 2002). Several meta-analyses comparing more than 20 studies conducted over the past two decades have made a case that silver alloy (but not silver oxide)-coated devices are effective in reducing overall catheterassociated urinary tract infection (CAUTI) rates by up to 45% (Davenport and Keeley, 2005; Slawson et al., 1992). However, it is important to note that many of those studies could not show significance statistically, and wide variations in overall efficacy were observed (Davenport and Keeley, 2005; Jacobsen et al., 2008). In addition, the higher cost of these devices over standard catheters further clouds the issue. Ultimately, despite intense research and over two decades of use, it currently remains unclear as to

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whether silver-coated urethral catheters are the best clinically viable tool for significantly decreasing CAUTI (Jacobsen et al., 2008). Antimicrobial hydrogels Maeyama et al. (2005) developed a novel surface based on a biodegradable polymer consisting of an antimicrobial catechin found in green tea extracts (Chosa et al., 1992). The surface-erodible polymer, epigallocatechin-3gallate (EGCg), showed a dose-dependent inhibitory effect on Escherichia coli biofilm formation and an enhanced dose-dependent effect on E. coli biofilm destruction. The antimicrobial activity of EGCg has since been shown to be due to DNA gyrase inhibition, the enzyme also targeted by the quinolone antibiotics (Gradisar et al., 2007). The novel, hydrophilic surface treatment swelled in the presence of a physiological fluid; subsequently releasing EGCg to the bulk phase. Over time, surface erosion proceeds and more EGCg is released resulting in high local concentrations of EGCg in near proximity to the surface. Such high concentrations contribute to the destruction of the three-dimensional biofilm interface as well as the biofilm– substrate interface (Maeyama et al., 2005). The direct consequences of these effects are the subsequent detachment, removal or washout of the biofilm. Further studies are required to determine whether these effects can be translated to the clinical situation. Numerous other antimicrobial factors have also been coated on to urological devices for biofilm prevention and treatment. These have involved a plethora of antibiotics – including gentamicin, cefazolin and several quinolones (DiTizio et al., 1998; John et al., 2007; Pugach et al., 1999) – as well as innate immune factors such as nitric oxide (Nablo et al., 2005). In addition, certain hydrogels have been developed that release stored antimicrobials only upon the application of ultrasonic waves to allow for tight regulation of compound release (Norris et al., 2005). Overall, these studies vary widely in their results, ranging from no perceived effect to complete biofilm prevention; the results also tend to be antimicrobial and strain specific. These largely conflicting results, combined with generally short elution periods observed with these types of coatings and the growing trend of antimicrobial resistance and tolerance, have prevented this line of research from being deemed a clinical success. Triclosan Triclosan is a broad-spectrum antimicrobial that has been in use for over 40 years and is currently found in numerous consumer and medical products – including toothpastes, surgical scrubs and vicryl suture (Bhargava and Leonard, 1996). At sub-lethal concentrations it acts by inhibiting

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the bacterial enoyl-acyl carrier protein (ACP) reductase (FabI), a critical enzyme in fatty acid biosynthesis, while at bacteriocidal concentrations it is believed to act through multiple, non-specific mechanisms including membrane destabilization (Heath et al., 1998; Heath et al., 2002). Two urological approaches that have been recently applied to investigate its ability to prevent device-associated biofilm formation and infection are: (a) elution from a ureteral stent impregnated with triclosan (Triumph®, Boston Scientific Corp. Inc., Natick, MA, USA) and (b) elution from catheter balloons filled with the antimicrobial (Jones et al., 2006). Triclosaneluting stents demonstrated significant antimicrobial and infection clearance effects during preliminary in vitro (Chew et al., 2006) and in vivo (Cadieux et al., 2006) studies, respectively, as well as anti-inflammatory effects in cell culture (Elwood et al., 2007). However, a clinical trial involving chronically stented patients showed no improvement in device colonization or urine culturing despite a significant reduction in antibiotic usage compared with controls (Cadieux et al., 2008). Thus, at least within this complicated patient population, the elution of triclosan alone from the device was unable to prevent colonization and biofilm formation. Jones et al. (2006) conducted preliminary in vitro experiments involving catheters whose retention balloons were filled with triclosan. Their results were very encouraging and are discussed in more detail in Stickler’s chapter on catheters (Chapter 7).

1.8.2 Biofilm-disrupting agents In this chapter we have already discussed cell–cell communication within biofilms as a means for microorganisms to control biofilm development and spread. However, these factors also offer a potential opportunity to influence biofilm development on medical devices either through surface coating or exogenous application (Rasmussen and Givskov, 2006). Studies have already examined several of these quorum-sensing factors on Pseudomonas aeruginosa biofilms and they have been shown to inhibit biofilm formation (Laux et al., 2002) and promote biofilm dissolution (Hentzer et al., 2002). In addition, ureteral stents coated with the quorum-sensing inhibitor RNA III inhibiting peptide from Staphylococcus aureus were found to resist biofilms of that organism in vitro (Balaban et al., 2005) and in a rat model of device infection, especially when combined with the antibiotic teicoplanin (Cirioni et al., 2007). Although still in its infancy, this line of research may prove to be fruitful as it specifically targets biofilm organisms by using their own naturally evolved signaling factors against them. Furthermore, these factors can be combined with other agents such as antimicrobials to either kill or inhibit the growth of organisms released from the biofilms before they can adhere again.

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1.8.3 Novel antifouling coatings Following many uses in industrial applications, diamond-like coatings (DLCs) are being adapted for biomedical applications. These amorphous carbon coatings are known for their superior tribologic properties and their biocompatibility. Laube et al. (2007) developed a low-temperature, lowpressure, plasma-enhanced chemical vapor deposition technique for coating polymeric medical implants with DLC. The coating was assessed for its ability to decrease the extent of crystalline biofilm formation, as well as the potential stent side effects and patient comfort. Results indicated that in vivo biofilm formation and encrustation were effectively reduced and patients were more tolerant of replacement of the low-friction devices. While the precise reasons for the positive performance characteristics of DLC are still not clear, results certainly warrant additional research. Previous work involving hydrogels containing poly(ethylene glycol) (PEG) have shown considerable success as antifouling agents due to PEG’s hydrophilic nature, high degree of mobility, and steric hindrance in its chemical structure. The polyether structure of PEG allows it to couple to large numbers of water molecules, reducing its coefficient of friction and driving it to behave more like a fluid. Importantly, PEG has shown considerable success in resisting not only bacterial attachment, but that of proteins and mammalian cells as well, suggesting that it may be able to resist conditioning film development. To date, the main concerns with using PEG have been the inability to anchor enough of the molecules to generate a robust, high-density coating and prevent its thermal, oxidative or hydrolytic degradation during this process. A novel approach for attaching PEG and other molecules to device surfaces is derived from nature, employing peptide mimics based upon the adhesive proteins involved in the adherence of mussels to underwater surfaces. In vitro and in vivo studies involving surfaces coated with PEG conjugated to one of these biological glues, namely 3,4-dihydroxyphenylalanine (DOPA) demonstrated the inhibition of both conditioning and biofilm formation, as well as a reduction in UPEC adherence within a rabbit model of cystitis. Future studies are required to address its true clinical potential.

1.8.4 Novel stent designs While urinary stenting has been used in medicine for over 2000 years (Laube et al., 2007), modern stents have changed very little in terms of overall design. Recent investigations (Tong et al., 2007) have brought to light an apparent lack of clear design rationale for some aspects of common stents; in particular, the number, size and orientation of the pass-through holes in the stent. Some researchers have even questioned the efficacy and

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necessity of the central bore associated with a typical stent. A novel stent designed for biliary drainage was developed with a winged cross-section (Raju et al., 2006). Numerical modeling confirmed that this design offered a larger surface area for flow, higher flow velocity and an increased flow rate compared with that of a conventional tubular stent. In addition, the lack of a central bore eliminates the potential for luminal occlusion from biofilms. Though originally designed for biliary drainage, such a novel design may be adapted for urological application. Responding to high stenting failure rates in patients with extrinsic ureteral obstruction, some minor modifications have been made to the traditional stent with limited success (Diaz-Lucas et al., 1997; Shick et al., 1998; Tliguri et al., 2000). Representing a more radical departure, some (Rotariu et al., 2001) attempted the simultaneous insertion of two ipsilateral stents yielding better flow rates with time. Unfortunately, the logistical challenges associated with installation and removal may outweigh the benefits of the configuration. In light of this, monolithic double-‘barreled’ or double-lumen stents were developed by Gyrus ACMI (Southborough, MA, USA) and Cook Urological (Bloomington, IN, USA). Hafron et al. (2006) evaluated these dual-lumen stents in ex vivo porcine kidney models with extrinsic ureteral obstruction. The prototype double stents provided excellent flow that rivaled that of two stents juxtaposed. While these stents were assessed under conditions of extrinsic obstruction, it could be rationally postulated that the built-in redundancy of the design may increase their efficacy under conditions of intrinsic blockage.

1.9

Future trends

Biofilm prevention and eradication have been at the forefront of both clinical and industrial research on urological devices for several decades, and countless studies have looked into developing biofilm-resistant materials, surface coatings, signaling molecules, antimicrobials and physical treatments with proven clinical benefit. Although many of these studies have shown powerful in vitro and even in vivo animal model evidence of biofilm inhibition, nothing has ultimately led to a device or treatment modality that would be considered a clinical triumph in the field. However, this is not entirely surprising as biofilm-associated device infections are multi-factorial and depend not only on bacterial aspects, but also on the host and environment as well. Exhaustive studies involving a wealth of antimicrobials have revealed that the vast majority of biofilms cannot be eradicated by these compounds once formed, and that their repeated use in these situations may instead promote resistance. That being said, their ability to at least slow biofilm formation and clear planktonic organisms is of critical importance and their use toward this end must be continued until better therapeutic

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options are available. Novel device designs and coatings are constantly under development and examination in attempts to build upon previous failures and achieve a resistant device. Currently, the most promising molecules are those involved in bacterial communication, termed ‘quorum sensing’, as they can target biofilm organisms specifically and may be able to break them down once formed. Their application appears thus far to be optimal when utilized in conjunction with antimicrobials, since planktonic organisms and those released from existing biofilms can be killed. Ultimately, success against biofilms and their related infections may not come in the form of complete prevention or clearance, but more in their clinically relevant control.

1.10

Conclusions

The development of biofilms by microorganisms is both a natural and vital part of their biological imperative, namely survival. The fact that these organisms also form biofilms on and within the human body can be rationalized as a simple extension of this natural tendency, not necessarily a distinct phenomenon. Biofilms pose an enormous challenge in medicine as they can form rapidly on almost any surface and once formed, are virtually impossible to remove. This is understandable since microorganisms use multiple strategies to form biofilms and can rapidly adjust to changing environmental conditions including resistance to host immune responses and antimicrobials. Most strategies to date for inhibiting biofilms deal with the use of antimicrobials but generally only work against planktonic organisms shed from the biofilm. It is interesting to note that the human immune system is largely successful against infectious microorganisms that enter the body through the concerted efforts of a wealth of innate and acquired immune strategies. Thus, it may simply be that one or two strategies employed at a time are not sufficient to prevent or remove biofilms and their related infections. Perhaps a multi-faceted approach to device development employing eluting antimicrobials with surface-bound silver and biofilm-disrupting agents may be more successful. This strategy has the potential not only to attack the pathogen in multiple ways simultaneously, but can also allow the host’s natural immune system to play a role as well. In addition, the use of oral and systemic antibiotics prior to device placement or between device changes to sterilize the tract can be critical to preventing infections.

1.11

Sources of further information and advice

In terms of general information on biofilms, including powerful figures and interesting commentary, The Biofilm Primer by Dr Bill Costerton is a great

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resource (Costerton, 2007a). Costerton is arguably one of the leading global experts on biofilms and has also written several solid reviews on the topic encompassing a multitude of different organisms, devices and medical conditions (Costerton et al., 1999; Costerton et al., 2005; Donlan and Costerton, 2002; Stoodley et al., 2002b). Antibiotic resistance and the development of persister cells are described in several excellent reviews by Lewis (Lewis, 2001; Lewis, 2005; Lewis, 2007). Information regarding urinary tract biomaterials, ureteral stent technology and catheters can be found in reviews by Beiko et al. (2004), Denstedt et al. (2000) and Carr (2000), respectively. Infections caused by biofilms of urease-producing organisms and the resulting encrustation of urinary tract devices are covered by Stickler in this textbook (Chapter 7) as well as in several reviews (Jacobsen et al., 2008; Morris et al., 1999; Stickler et al., 1998). Hydrodynamics and their impact on biofilm formation and device failure are covered in reviews by Nguyen et al. (2005), Purevdorj et al. (2002) and Stoodley et al. (1994). A comprehensive book entitled Medical Biofilms, edited by Jass, Surman and Walker covers a wealth of diverse topics concerning all forms of clinically relevant biofilms (Jass et al., 2003).

1.12

References

abraham, w. r. (2006) Controlling biofilms of gram-positive pathogenic bacteria. Curr Med Chem, 13, 1509–24. allesen-holm, m., barken, k. b., yang, l., klausen, m., webb, j. s., kjelleberg, s., molin, s., givskov, m. and tolker-nielsen, t. (2006) A characterization of DNA release in Pseudomonas aeruginosa cultures and biofilms. Mol Microbiol, 59, 1114–28. allison, d. g. and gilbert, p. (1995) Modification by surface association of antimicrobial susceptibility of bacterial populations. J Ind Microbiol, 15, 311–7. anderl, j. n., franklin, m. j. and stewart, p. s. (2000) Role of antibiotic penetration limitation in Klebsiella pneumoniae biofilm resistance to ampicillin and ciprofloxacin. Antimicrob Agents Chemother, 44, 1818–24. anderson, g. g., dodson, k. w., hooton, t. m. and hultgren, s. j. (2004) Intracellular bacterial communities of uropathogenic Escherichia coli in urinary tract pathogenesis. Trends Microbiol, 12, 424–30. anwar, h., dasgupta, m., lam, k. and costerton, j. w. (1989) Tobramycin resistance of mucoid Pseudomonas aeruginosa biofilm grown under iron limitation. J Antimicrob Chemother, 24, 647–55. balaban, n., stoodley, p., fux, c. a., wilson, s., costerton, j. w. and dell’acqua, g. (2005) Prevention of staphylococcal biofilm-associated infections by the quorum sensing inhibitor RIP. Clin Orthop Relat Res, 437, 48–54. beiko, d. t., knudsen, b. e., watterson, j. d., cadieux, p. a., reid, g. and denstedt, j. d. (2004) Urinary tract biomaterials. J Urol, 171, 2438–44. beyenal, h. and lewandowski, z. (2002) Internal and external mass transfer in biofilms grown at various flow velocities. Biotechnol Prog, 18, 55–61.

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shah, d., zhang, z., khodursky, a., kaldalu, n., kurg, k. and lewis, k. (2006) Persisters: a distinct physiological state of E. coli. BMC Microbiol, 6, 53. shick, r., seidal, e. m., kalem, t., volkmer, b. and planz, k. (1998) New endoureteral double-J stent resists extrinsic ureteral compression. J Endourol, 12, 37–40. silverstein, a. and donatucci, c. f. (2003) Bacterial biofilms and implantable prosthetic devices. Int J Impot Res, 15, Suppl 5, S150–4. silverstein, a. d., henry, g. d., evans, b., pasmore, m., simmons, c. j. and donatucci, c. f. (2006) Biofilm formation on clinically noninfected penile prostheses. J Urol, 176, 1008–11. slawson, r. m., van dyke, m. i., lee, h. and trevors, j. t. (1992) Germanium and silver resistance, accumulation, and toxicity in microorganisms. Plasmid, 27, 72–9. sobel, j. d. (1990) Bacterial vaginosis. Br J Clin Pract Suppl, 71, 65–9. stamey, t. a., timothy, m., millar, m. and mihara, g. (1971) Recurrent urinary infections in adult women. The role of introital enterobacteria. Calif Med, 115, 1–19. stapleton, a. e., fennell, c. l., coder, d. m., wobbe, c. l., roberts, p. l. and stamm, w. e. (2002) Precise and rapid assessment of Escherichia coli adherence to vaginal epithelial cells by flow cytometry. Cytometry, 50, 31–7. stewart, p. s. and costerton, j. w. (2001) Antibiotic resistance of bacteria in biofilms. Lancet, 358, 3. stickler, d. (1996) Bacterial biofilms and the encrustation of urethral catheters. Biofouling, 94, 293–305. stickler, d. (1999) Biofilms. Curr Opin Microbiol, 2, 270–5. stickler, d., ganderton, l., king, j., nettleton, j. and winters, c. (1993) Proteus mirabilis biofilms and the encrustation of urethral catheters. Urol Res, 21, 407–11. stickler, d., morris, n., moreno, m. c. and sabbuba, n. (1998) Studies on the formation of crystalline bacterial biofilms on urethral catheters. Eur J Clin Microbiol Infect Dis, 17, 649–52. stoodley, p., cargo, r., rupp, c. j., wilson, s. and klapper, i. (2002a) Biofilm material properties as related to shear-induced deformation and detachment phenomena. J Ind Microbiol Biotechnol, 29, 361–7. stoodley, p., debeer, d. and lewandowski, z. (1994) Liquid flow in biofilm systems. Appl Environ Microbiol, 60, 2711–16. stoodley, p., hall-stoodley, l. and lappin-scott, h. m. (2001a) Detachment, surface migration, and other dynamic behavior in bacterial biofilms revealed by digital time-lapse imaging. Methods Enzymol, 337, 306–19. stoodley, p., lewandowski, z., boyle, j. d. and lappin-scott, h. m. (1999a) The formation of migratory ripples in a mixed species bacterial biofilm growing in turbulent flow. Environ Microbiol, 1, 447–55. stoodley, p., lewandowski, z., boyle, j. d. and lappin-scott, h. m. (1999b) Structural deformation of bacterial biofilms caused by short-term fluctuations in fluid shear: an in situ investigation of biofilm rheology. Biotechnol Bioengng, 65, 83–92. stoodley, p., sauer, k., davies, d. g. and costerton, j. w. (2002b) Biofilms as complex differentiated communities. Annu Rev Microbiol, 56, 187–209. stoodley, p., wilson, s., hall-stoodley, l., boyle, j. d., lappin-scott, h. m. and costerton, j. w. (2001b) Growth and detachment of cell clusters from mature mixed-species biofilms. Appl Environ Microbiol, 67, 5608–13.

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stramer, s. l. and starzyk, m. j. (1978) Improved growth of Thermus aquaticus on cellular lysates. Microbios, 23, 193–8. swidsinski, a., mendling, w., loening-baucke, v., ladhoff, a., swidsinski, s., hale, l. p. and lochs, h. (2005) Adherent biofilms in bacterial vaginosis. Obstet Gynecol, 106, 1013–23. tack, k. j. and sabath, l. d. (1985) Increased minimum inhibitory concentrations with anaerobiasis for tobramycin, gentamicin, and amikacin, compared to latamoxef, piperacillin, chloramphenicol, and clindamycin. Chemotherapy, 31, 204–10. tliguri, m., nouri, m., you, r., habb, f., gattegno, b. and thibault, p. (2000) Value of cloverleaf double J ureteral stents in the treatment of extrinsic ureteral compression. Progres en Urologie, 10, 92–4. tong, j. c. k., sparrow, e. m. and abraham, j. p. (2007) Numerical simulation of the urine flow in a stented ureter. J Biomech Eng, 129, 6. towler, b., cunningham, a., stoodley, p. and mckittrick, l. (2007) A model of fluid–biofilm interaction using a Burger Material law. Biotechnol Bioengng, 96, 13. vahidi, b. and fatouraee, n. (2007) A numerical simulation of peristaltic motion in the ureter using fluid structure interaction. In Proceedings of the 29th Annual International Conference of the IEEE EMBS Cite Internationale, IEEE Lyon, France, pp. 1168–71. veeh, r. h., shirtliff, m. e., petik, j. r., flood, j. a., davis, c. c., seymour, j. l., hansmann, m. a., kerr, k. m., pasmore, m. e. and costerton, j. w. (2003) Detection of Staphylococcus aureus biofilm on tampons and menses components. J Infect Dis, 188, 519–30. vigeant, m., rothstein, s. n. and cacciatore, j. j. (2005) Removal of adhered bacteria by surfactant and shear. In AIChE Annual Meeting, Conference Proceedings, American Institute of Chemical Engineers, Cincinnati, Ohio, 528f. vuong, c., voyich, j. m., fischer, e. r., braughton, k. r., whitney, a. r., deleo, f. r. and otto, m. (2004) Polysaccharide intercellular adhesin (PIA) protects Staphylococcus epidermidis against major components of the human innate immune system. Cell Microbiol, 6, 269–75. whitchurch, c. b., tolker-nielsen, t., ragas, p. c. and mattick, j. s. (2002) Extracellular DNA required for bacterial biofilm formation. Science, 295, 1487. wilson, s. k. and delk, j. r., ii (1995) Inflatable penile implant infection: predisposing factors and treatment suggestions. J Urol, 153, 659–61. wright, k. j., seed, p. c. and hultgren, s. j. (2007) Development of intracellular bacterial communities of uropathogenic Escherichia coli depends on type 1 pili. Cell Microbiol, 9, 2230–41. yoshinari, m., kato, t., matsuzaka, k., hayakawa, t., inoue, t., oda, y., okuda, k. and shimono, m. (2006) Adsorption behavior of antimicrobial peptide histatin 5 on PMMA. J Biomed Mater Res B Appl Biomater, 77, 47–54. ziebuhr, w., krimmer, v., rachid, s., lossner, i., gotz, f. and hacker, j. (1999) A novel mechanism of phase variation of virulence in Staphylococcus epidermidis: evidence for control of the polysaccharide intercellular adhesin synthesis by alternating insertion and excision of the insertion sequence element IS256. Mol Microbiol, 32, 345–56.

2 In vivo models for ureteral stents M. K. L O U I E, A. J. G A M B OA and R. V. C L AY M A N, UCI Medical Center, USA

Abstract: Animal models have long been used as a surrogate for human testing for the creation of innovative new medical therapies, from chemotherapies to new surgical instruments. The field of endourology is no exception with rats, rabbits, dogs, and pigs having been used most commonly to investigate new devices, procedures, or materials. This chapter will explore the history of in vivo studies of ureteral stents, compare the benefits of the in vivo model with the in vitro model, examine each animal’s particular role and contribution to the study of ureteral stents, and finally provide some perspective on future studies utilizing in vivo animal models. Key words: in vivo animal model, ureteral stent, rat model, rabbit model, dog model, pig model, endourology model.

2.1

Introduction

Animal models have long been used as a surrogate for human testing for the creation of innovative new medical therapies from chemotherapies to new surgical instruments. Experimental animals will remain a critical tool for bridging basic sciences research via translational studies leading to improvements in health care for humans and animals. Numerous factors enter into the selection of the appropriate animal for model studies. The field of endourology is no exception with rats, rabbits, dogs, and pigs having been used most commonly to investigate new devices, procedures, or materials. This chapter will explore the history of in vivo studies of ureteral stents, compare the benefits of the in vivo model with the in vitro model, examine each animal’s particular role and contribution to the study of ureteral stents, and finally provide some perspective on future studies utilizing in vivo animal models.

2.1.1 History of in vivo models for ureteral stents The modern indwelling ureteral stent was first described by Zimskind et al. in 1967; they reported the cystoscopic insertion of a silicone rubber tube 42

In vivo models for ureteral stents

43

into the ureter to treat stricture disease, malignant obstruction, and ureterovaginal fistula. Almost a decade later, Gibbons et al. (1976) designed a stent with a radiopaque tip, flared wings, spring coil reinforcement, and a distal flange to resolve the issues of radiolucency, downward expulsion, external compression, and upward migration. The creation of double-J and double pigtail stents would come a few years later when Finney (1978) and Hepperlen et al. (1978) respectively reported their development. Only after their introduction in humans were further studies in animals performed to characterize these devices. To date, there are only a handful of studies describing the accuracy of certain animal models for the in vivo study of ureteral stent properties. Most of these are based on anatomical considerations while other aspects such as urine biochemistry, renal physiology, and immunological characteristics are not fully addressed.

2.1.2 In vitro versus in vivo ureteral stent models In vitro models for ureteral stents have been used to a wide extent due to many factors – including decreased cost, increased control of external variables, reduced complexity of experimental equipment and laboratory requirements, and ethical concerns regarding the use of animals. However, in vitro models do not address many factors that influence the performance of a ureteral stent. Besides the obvious lack of a dynamic system involving interplay between urine, urothelium, ureteral peristalsis, stent reflux, and bladder filling and emptying, there is a lack of immunoreactivity in the in vitro setting. Tunney et al. (1996) developed an in vitro model for biomaterial encrustation consisting of portions of ureteral stents suspended in artificial urine, incubated at 37 °C, and the atmosphere equilibrated to 5% CO2 with stir bars constantly mixing the solution. They also performed this experiment in collected human urine with and without stone pieces crushed into the urine. The results showed similar encrustation composition to that found in ureteral stents extracted from human patients. However, in the experimental arm without stone pieces crushed into the human urine there was no encrustation. The benefits of such an experimental system were the ability to control many of the external variables of encrustation and the ability to have high throughput testing of many biomaterials in a short period of time. Unfortunately, the model can only show one result if the artificial urine was supersaturated with the exact recipe for calcium phosphate stones and the human urine was also given crushed calcium phosphate stones. The authors have since modified their experimental apparatus to include a continuousflow peristaltic pump to simulate the flow of urine (Gorman et al., 2003). This example illustrates the difficulty of studying a ureteral stent in vitro as many factors influence the biomaterial to form a biofilm or to encrust.

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Biofilm formation has been linked to direct bacterial binding to biomaterials, but more likely it is a two-fold process. Initially a conditioning film is produced by host urinary substances that adhere to the surface of the biomaterial within minutes of placement. This film includes proteins, electrolytes, and other organic materials. Secondly, this film changes the biomaterial surface to ease bacterial adhesion and subsequently colonization. Encrustation can occur in both infected and sterile urine. The mechanism of encrustation in infected urine is that of urine alkalinization due to urease-producing organisms which form magnesium ammonium phosphate and calcium apatite stones (Denstedt et al., 2000). In sterile urine, the mechanism by which encrustation occurs is not completely elucidated. A study by Keane et al. (1994) randomly selected 40 patients with indwelling stents placed for a variety of indications including stones, ureteropelvic junction obstruction, renal transplant, retroperitoneal fibrosis, and upper tract urothelial carcinoma among others. The stents were polyurethrane in all cases except for one silicone stent. All patients received one 80 mg dose of i.v. gentamicin at the time of stent insertion. The duration the stents were indwelling was 2–36 weeks (median 12 weeks). At the time of removal, the stents were sectioned for microscopic analysis and microbiological identification. Their results showed that a viable microbial biofilm (>104 colony forming units (CFU)/cm3 of stent) was apparent in 11 (28%) of the stents as shown by scanning electron microscopy. Thirteen different organisms were identified on these stents. Encrustation (microscopic or macroscopic) was seen in 23 (58%) stents, but only 8 of these stents had a biofilm associated. Encrustation was heaviest in the stents removed after 6 weeks, with 75% of stents left indwelling longer than 23 weeks developing encrustation. Of the 18 patients with a stent placed for an obstructive stone, 15 developed encrustation of the stent, but only 6 had biofilms associated. This small study showed that the presence of urolithiasis was a major determining factor in stent encrustation, but was unable to link biofilm formation with encrustation. For an indwelling stent, a biofilm or encrustation (or both or neither) may occur; as yet, the precise reason for each state has not been elucidated. The ideal in vivo model would be anatomically identical to humans with renal, ureteral, and bladder physiology as similar as possible. Urine biochemical composition and its variation with fluid/food intake as well as immunoreactivity of the tissue when in contact with the stent would also need to be established. As such an in vitro model does not yet exist, in vivo models are necessary to expedite the transfer of experimental results to clinical application. Only with testing in the in vivo animal model can a ureteral stent undergo all of the possible interactions that might occur in the clinical setting, bearing in mind that each animal’s particular urinary system traits differ in some way from the human situation.

In vivo models for ureteral stents

2.2

45

Commonly used animal models

2.2.1 Rat The rat urinary system is seldom used as an in vivo model for ureteral stents due to the obvious size difference and therefore, difficulty with endourological manipulation. However, much of what we know about mammalian renal physiology comes from studies on the rat. In fact, rat renal physiology is very similar to that of the human, including the production of hormones and urine biochemistry. The rat kidney is a paired organ system in the retroperitoneal space in the posterior part of the abdomen. Each adult rat kidney weighs approximately 0.8–1.4 g. Gross anatomical structure is similar to that of the human kidney with the concave side of the kidneys carrying the hilum through which the vessels, nerves, and lymphatics pass into and out of the kidney. The urine drains from a single renal papilla and is carried down into the renal pelvis into the ureter. Wolf et al. (1996) performed a comparison of the ureteral microanatomy of the rat and that of other animal models. Mean cross-sectional thicknesses of the ureteral layers were very significantly different in the rat and the human ureter. The relative percentage of each ureteral layer (i.e. 10% epithelium, 30% lamina propria, etc.) was also significantly different. The earliest use of the rat model for studying ureteral stents was for renal transplantation studies (Frodin and Engberg, 1975; Jablonski et al., 1995). Most of these studies concerned the healing of the ureter over a ureteral splint and the different techniques for ureteral anastomosis. Hildebrandt et al. (2001) tested heparin-coated stents by open incision of the ureter to implant the stents. Studies on rat renal physiology revealed that the urine biochemistry in rats was similar to humans. Miyake et al. (1998) characterized the expression of Tamm–Horsfall protein in stone-forming rat models. Atmani and colleagues looked at inhibitors of calcium oxalate crystallization isolated in rat urine (Atmani and Khan, 1995; Atmani et al., 1996). The rat kidney approximate glomerular filtration rate is 0.9–1.2 mL/min/per gram of kidney weight. Normal daily urine output for the rat kidney is 15–30 mL. Table 2.1 shows the comparative urinary biochemistry of the rat and other animals, including values for humans. Other studies have subsequently justified the use of the rat bladder as a viable model of in vivo stent encrustation and biofilm formation. One rat bladder model involves the implantation of small pieces of the various biomaterials under study into the rat bladder; these pieces are then harvested after an incubation period for evaluation (Yasuda et al., 1999; Cirioni et al., 2007; Orlando et al., 2008). Minardi and colleagues (2007)

0.05–0.09

0.3–0.5

0.2–3.2

0.6–1.4

Rabbit

Dog

Swine

Human

Specific gravity

Sodium (mmol/L)

5–8

Citrate (mg/dL)

Uric acid (mg/dL)

2.5 ± 1.2 5.2 ± 0.9 2.9 ± 2.6

33.6 ± 28.3 89 ± 56 66.1 ± 22.1 52.3 ± 36.3

4 ± 5.6

1.92 ± 1.15

6.8 ± 3.9

2.6 ± 1.7

3.3 ± 1.3

Phosphorus (mg/dL)

7.3 ± 3.6

13.5 ± 13.7

3±2

61.6 ± 27

38.2 ± 29.9

102 ± 58

22.6 ± 21

0.185 77.4 mmol/24 h

Magnesium (mg/dL)

176 ± 200.6 47.8 ± 27.6

2.8

Calcium (mg/dL)

45.1 ± 20.5 11.7 ± 6.1

4.5 ± 4.7

10.3 ± 2.9

5.9 ± 3.4

0.0058 0.09 28 mmol/24 h mmol/24 h

Oxalate (mg/dL)

1.016–1.035 108.7 ± 43.4 2.3 ± 0.9

5.8–6.7 1.003–1.025

6.0–7.0 1.001–1.060

7.6–8.8 1.003–1.036

7.3–8.3 1.050–1.062 200

pH

Note: data are approximate and vary significantly from strain to strain, and depend on age, sex, and diet. Adapted from G. N. Box, J. B. Abraham, H. J. Lee, L. A. Deane, C. S. Abdelshehid, D. R. Tyson, E. R. Elchico, R. Alipanah, J. F. Borin, J. R. Asplin, E. M. McDougall, and R. V. Clayman (unpublished data, 2007); Loeb and Quimby (1999); Suckow et al. (2006).

0.01–0.02

Rat

Urine flow (mL/min)

Table 2.1 Comparative normal urine biochemistry of the rat, rabbit, dog, pig, and human

In vivo models for ureteral stents

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implanted infected ureteral stents into subcutaneous pockets made in their rat model. The advantages of the rat model include a similar biochemistry to human urine, decreased procedural expenses compared with other animal models, and the ease of acquiring the animals. The disadvantage is that the urinary tract anatomy is not only much smaller than the human urinary tract, but it is also unipapillary.

2.2.2 Rabbit The rabbit has been used in almost all areas of biomedical research and has been a significant contributor in many specialized areas of scientific investigation. The rabbit kidneys are paired organs, lying in the retroperitoneum, approximately in the position of the embryonic intermediate cell mass from which they are formed. During development, one kidney is often displaced more than the other by the pressure of adjacent organs so that the symmetrical disposition of the pair is destroyed. In the human, the right kidney is situated lower than the left on account of the pressure of the right lobe of the liver. In the rabbit, on the other hand, the left kidney is displaced further back than the right by the posterior expansion of the greater curvature of the stomach. The right kidney is placed a little farther forward than the left and is largely covered by the right posterior lobule of the liver (Barone, 1973). The rabbit kidneys appear as solid organs, brownish in color and bean-like in general shape, enclosed with a fibrous coat. It appears as an almost continuous mass; however, slight traces of lobulation can be identified. In the rabbit, there is only a single renal papilla and the expanded end of the ureter, the renal pelvis, is thus undivided (Bensley, 1948). In the unipyramidal kidney of the rabbit, the renal pelvic tissues form two septa, which enclose the medullary pyramid. The pelvic cavity extends up between the septa and the pyramid to the fornix and then turns outwards to form a series of secondary pouches on the outer surface of the pelvic septa (Sheehan and Davis, 1959). The rabbit ureter, like the rat ureter, is significantly different from the human ureter in terms of ureteral layer cross-sectional thicknesses and in the relative percentages of each of these layers (Wolf et al., 1996). The human ureter receives blood mainly from three vascular territories: the upper ureter is fed by branches from the renal arteries, the middle ureter is fed by branches from the common iliac or middle ureteric artery, and the lower ureter by branches from the vesicular or uterine arteries (Daniel and Shackman, 1952). Douglas and Hossler (1995) showed that the rabbit ureter is similar, although not identical, to the human ureter with respect to the vascular supply. The rabbit ureteral blood supply is dependent almost exclusively on a single set of longitudinal vessels running the full length of the organ. Proximally, blood is

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Biomaterials and tissue engineering in urology

supplied to each ureter by single branches from the renal artery and vein, and distally by single branches of the vesicular artery and vein. The only other vascular connections to the ureter include continuities with the capillary beds of the kidney pelvis cranially and the bladder caudally. The consequence of such an arrangement is that most of the length of the ureter is completely dependent for nutrition on the integrity of the longitudinal adventitial vessels. The rabbit urinary bladder is a muscular sac that lies in the ventral posterior portion of the rabbit’s abdominal cavity. Its normal function includes filling and emptying, accompanied by distention and relaxation. The urinary bladder requires a rich blood supply to maintain its functions. The bladder is supplied with blood by single, left, and right vesicular branches of the internal or external iliac arteries. The serpentine vesicular arteries extend along the lateral borders of the bladder from base to apex, just deep to the serosal surface, and send dorsal and ventral branches to supply the dorsal and ventral bladder wall. Veins accompany the arteries and exhibit numerous valves. These specializations of the vasculature appear to enhance blood flow in the bladder wall during distention (Hossler and Monson, 1995; Hossler and Monson, 1998). Table 2.1 shows the comparative urine biochemistry of the rabbit and other commonly used animal models, and also human urine composition. The normal daily urine production of the rabbit is about 20–35 mL/kg per 24 h. The average pH is very alkaline (range 7.6–8.8). The specific gravity ranges from 1.003 to 1.036. The latter is difficult to measure due to the presence of urine crystals, ammonium magnesium phosphate, and calcium carbonate (Mitruka and Rawnsley, 1977). The presence of minerals can vary the urine’s appearance from a translucent yellow color with a water-like consistency to a reddish brown thick creamy paste (Cheeke and Amberg, 1973). The high pH favors the development of bacteria and urinary tract infection. The alkaline urine has also been reported to predispose to precipitation of calcium carbonate crystals in urine (Lee et al., 1978). The rabbit has a unique calcium metabolism (Cheeke and Amberg, 1973; Kamphues, 1991). Calcium is used for a variety of processes, the most common being the maintenance of bone and muscle. Humans and most other domestic animals tend to absorb calcium out of the diet in proportion to what their body needs at that particular time. The calcium that is not needed by the body is passed into the gastrointestinal tract, primarily through bile, and is excreted in the feces. Rabbits, however, have a different method of dealing with digestible calcium in the diet. They tend to absorb calcium in direct proportion to the digestible calcium that is in the diet whether or not their body needs that extra amount at that moment. Therefore, high blood calcium concentrations (3.0–4.0 mmol/L) are a common finding after a calcium-rich diet, such as alfalfa. Rabbits then excrete the excess calcium

In vivo models for ureteral stents

49

they do not need, primarily through the kidneys. The calcium is excreted in the urine in the form of calcium carbonate. This potentially predisposes them to urolithiasis (Buss and Bourdeau, 1984). Although the exact mechanism of stone formation is poorly defined, a high-calcium diet and urinary tract infection have been widely implicated. Other contributing factors include metabolic disorders, nutritional imbalance, and low water intake (Harkness and Wagner, 1989). The diagnosis of calciuria and urolithiasis in rabbits is made by combination of urinalysis, abdominal radiographs, stone analysis, urine cultures, and blood tests. Urine for urinalysis and bacterial culture is best collected into a sterile syringe using a urinary catheter or by ‘cystocentesis’ to obtain urine directly from the bladder, without the potential contamination from the distal urethra or genital tract. Urine collection can also be done using metabolic cages providing some normal behavior to the rabbit, but this method of collection is prone to contamination from feces and food. Radiographs can provide information as to the location of calculi. The use of the laboratory rabbit has advantages that have made it a popular in vivo model. Its intermediate body and skeletal size, ear vascularization, immunological response to antigens, timed ovulatory response, nest-building phenomena, response to presumptive teratogenic drugs, and the simplicity of care relative to large animals have given the rabbit a unique importance in biomedical research. The rabbit is large enough to provide adequate quantities of tissues for experimental work without the pooling of samples, but is small enough to be more economical than the dog or monkey (Banks, 1989). Repeated blood samples for clinical chemistry and hematological determinations are easily obtained from the marginal ear vein or even the medial artery by a simple drip method, vacuum ear bleeding, or vacuum blood collection tubes (Hoppe et al., 1969; Manning et al., 1994). Two main disadvantages of the use of the rabbit in the laboratory include the potential of rabbit colonies to act as virtual storehouses of various diseases, and the rabbit’s extremely variable responses to most general anesthetics (Banks, 1989). In addition, its proclivity to encrust any foreign body placed in the urinary tract makes it a particularly severe, and possibly unfair, test subject for ureteral stents. However, the harsh rabbit bladder environment makes it the ideal litmus test for a urinary biomaterial attempting to claim encrustation resistance. The current authors (unpublished observations, 2008) used an in vivo rabbit bladder model to test the encrustation resistance of potential urinary tract biomaterials such as titanium, stainless steel, silicone, specially formulated silicones, specialty metals, and silicone nano-coated with titanium oxide. Results showed that all metals and metal-coated silicones encrusted in just 3 days after implantation, while the silicone materials did not encrust until 1 week. Biocompatibility testing was also performed in vitro using bladder epithelial cells and

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Biomaterials and tissue engineering in urology

showed again that the metal materials caused immediate cell death (current authors, unpublished data, 2008).

2.2.3 Dog The canine experimental model has been extensively used in urological research as it is one of the most studied research animals. Initially the gold standard for most urological anatomical and physiological studies, the canine model has been supplanted by the porcine model due in part to social mores regarding the use of canines in research in the past 20 years but also to the greater similarity between the porcine and the human kidney. The right canine kidney lies slightly cranial to the left because the stomach has evolved to lie on the left side of the abdomen, pushing the left kidney out of position. Lying close to the cranial pole of each kidney are the adrenal glands and inferiorly the ovaries in the female. Canine and human kidneys are similar in external configuration. However, unlike the human kidney, the canine kidney is unipapillate. Therefore, its entire internal architecture differs markedly from that of the human organ. The renal pelvis of the kidney is analogous to a minor calyx, but is called upon to drain an entire kidney. In order to do so, the renal pelvis conforms to the single medulla of this organ (Brodsky et al., 1977; Evans et al., 1979). Gross anatomic differences in renal vasculature exist between man and dog. Knowledge of intrarenal artery anatomy is important for performing intrarenal surgeries, which provide minimal blood loss and minimal injury to adjacent parenchyma. The vascular distribution in the human kidney is segmental but asymmetrical so that the functionally avascular plane lies about 0.5–1 cm posterior to Brodel’s white line (Smith and Boyce, 1968). By comparison, there is perfect anteroposterior symmetry in the arterial distribution in the dog. Consequently, midline bivalve nephrotomy in the human holds greater chance of transecting arcuate and interlobular arteries than it does in the dog (Klapproth, 1959; Ferber et al., 1966; Fuller and Huelke, 1973). Recently, Marques-Sampaio et al. (2007) showed that the anatomical relationship between arteries and the collecting system in the cranial pole of the dog is similar to that in man. This fact supports the use of the dog as an animal model for urological procedures at the cranial pole. In contrast, the ureteropelvic junction and the inferior pole of the kidney were not considered a good model for urological procedures because of the dissimilarity with the human anatomy (Sampaio and Favorito, 1993). In the dog, neither the ureter nor the bladder are similar to the human ureter and bladder. The dog ureter microanatomy differs significantly from the human ureter in cross-sectional thickness, size, relative ureteral layer thickness, and surface area (Wolf et al., 1996). The canine urinary bladder is a pear-shaped hollow organ. The rounded end points cranially, while the

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narrow end or neck points caudally and usually lies within the pelvic cavity. The main blood supply comes from the umbilical (becomes the cranial vesical artery) and the urogenital artery (visceral branch of the internal iliac artery) (Gordon, 1960). Table 2.1 shows the comparative urine biochemistry of the dog and the other commonly used animal models, and also normal human values. The normal daily production of urine for the dog is about 20–167 mL/kg per 24 h. The specific gravity ranges from 1.001 to 1.060. The pH is acidic at 6.0–7.0 with an osmolality of 176–1292 mOsm/kg (Mitruka and Rawnsley, 1977). The volume of urine secreted varies with the diet, fluid intake, climate, and activity. A high-protein diet increases the output of urine (National Research Council (US) et al., 1956). Puppies and younger dogs excrete proportionately larger volumes of urine per weight basis. The quantity of urine is of chief significance in 12- to 24-h specimens and should be interpreted in relation to the specific gravity. It is known that the dog can concentrate urine considerably more than man. Anatomical differences suggest physiological differences. Pfeiffer (1968) has postulated that this ability to hyperconcentrate urine rests in the anatomic relation of the pelvis to the medulla. In the dog, there is greater interface between both inner and outer medullary parenchyma and pelvic urine attributable to the large papilla that fills the renal pelvis as well as the pelvic diverticula that extend to the outer zone of the medulla, essentially surrounding its lobes with urine on three sides. This allows for a prolonged exchange of water and solutes relative to the human kidney where only the papillary tip has a parenchymal interface. The color of normal urine is due to certain pigments, the most important of which is urochrome. Normally, the color ranges from yellow to light amber in the dog. A cloudy urine or turbid urine is not necessarily pathological and may arise from vaginal or preputial contamination; otherwise cloudiness is due to precipitated salts, purulence, bacteria, epithelium, blood, and/or mucus. Decomposed urine develops an ammonia odor, while a similar odor in a freshly voided specimen suggests cystitis. Bacteriuria may likewise produce a foul odor. The normal dog urine is usually acidic. The composition of the dog’s food is perhaps the most important factor in determining the urine pH. Generally, foods of animal origin cause an acid tide while foods of vegetable origin cause alkaline urine. All urine on standing becomes alkaline because of ammonia formation resulting from decomposition of urea. In normal animals, the specific gravity varies from 1.018 to 1.060. In general, it depends on the amount of dissolved solids and varies inversely with the quantity of urine. Urinary calculi are a common and recurrent problem in dogs and cats (Bartges et al., 1999). Effective long-term management of urolithiasis depends on identification and manipulation of factors contributing to stone

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formation. The composition of stones in the dog, in descending order of frequency, is: struvite, calcium oxalate, urate cystine, xanthine, and silica. Urate calculi are more frequent with Dalmatians given that they share with humans the absence of the uricase enzyme. A strong association has been demonstrated in dogs between female sex and an increase risk of struvitecontaining uroliths (Ling et al., 1998; Houston et al., 2004). Most struvite stones in dogs are infection-induced and female dogs are at greatest risk for this condition. This is likely due, at least in part, to the anatomy of the female urethra which is short and wide compared with that of the male (Seaman and Bartges, 2001). Owing to the dog’s relatively large body size and friendly behavior (compared with other research animals), the canine has been used for functional studies more extensively than other laboratory animal species. Its convenience as a laboratory animal makes it an ideal experimental subject for urinary tract investigation; however, its unipapillate kidney and its overall favored status in society have largely precluded its ongoing use in most laboratories, and thus it has not been used commonly for ureteral stent studies.

2.2.4 Swine The porcine model is currently recognized as the most accurate comparative model for human renal anatomy. Unlike the rat, rabbit, and dog which have unipapillate kidneys, the pig has a true multipapillate system with associated minor and major calyces. Additionally, the larger size of certain pig strains emulates closely the renal size and ureteral length of humans. Evan et al. (1996) investigated the branching pattern of the porcine renal artery based on the initial work of Boyce et al. (1979). The work by Evan and colleagues was primarily used to support their use of the pig as a model for renal ischemic damage from shock wave lithotripsy. However, it was not until Sampaio and colleagues from Brazil (Sampaio et al., 1998) systematically studied the pig kidney anatomy that the porcine model was confirmed as the most representative urological model for the human kidney. In their seminal work, Sampaio et al. (1998) examined the porcine collecting system as it pertains to endourology. Fifty pigs weighing 60–80 kg had polyester resin injected into their ureter to fill the collecting system. These casts were then harvested and exposed to an acid bath. The results of the study showed that the dimensions of the pig kidney were similar to human kidneys; the mean weight of the pig kidney was 98 g, mean length of the pig kidney was 11.8 cm, mean cranial pole width was 5.64 cm, mean caudal pole width was 5.35 cm, and the mean thickness was 2.76 cm. The casts of the collecting system were used to characterize the pig pelviocaliceal system into two groups: Group A showed that the drainage of the

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mid-pole was dependent on cranial or caudal calyceal drainage; Group B showed mid-pole drainage that was independent of the polar calyces. Using this classification, Sampaio et al. found that 40% of pig kidneys exhibited Group A drainage while 60% exhibited Group B drainage; this compared to 62% of human kidneys with Group A drainage and 38% with Group B drainage. A perpendicular minor calyx draining directly into the renal pelvis or into a major calyx was seen in 18% of the pig kidneys compared with 11.4% in humans (Sampaio and Aragao, 1990). The authors concluded that the pig kidney is a good model for urological research owing to the similar morphometry of the pig and human kidney. Wolf et al. (1996) performed a comparative analysis of the microanatomy of the farm pig, minipig, human ureter, and several other animals (including the rat, rabbit, dog, and sheep). Surprisingly, the sheep ureter was the closest match to the human ureter based on relative ureteral layer sizes and ureteral layer content measurements, but was 45% smaller than the human ureter. In terms of ureteral cross-sectional thickness, the farm pig and the minipig were similar to the human ureter, but only the farm pig had the same cross-sectional surface areas. The minipig ureter was noted to be slightly larger than the human ureter. Overall, the rabbit and rat ureters differed in four of five relative measurments, and the dog differed in all five. The authors concluded that for endourological procedures and upper tract reconstruction, the farm pig and minipig were the least different in terms of the size and anatomy of the ureter, and also benefited due to the multipapillate nature of their kidneys. The Sampaio group also looked at both the intrarenal arterial and venous arrangements of the pig kidney. They found that in pigs, a single renal artery was found in all cases; while in humans, multiple renal arteries may be found in up to 30% of patients. Primary division of the renal artery was also different in that the porcine renal artery divides into a cranial and caudal branch in 93.4% of the observed cases, while in humans the renal artery branches into an anterior and posterior branch. Clinically this was an important observation because the porcine kidney avascular plane was transverse as opposed to being longitudinal in the human or dog kidney. Table 2.1 shows the comparative urine biochemistry of the Yorkshire swine and the other animal models, as well as normal human values. Standard clinical parameters of pigs are dependent on the strain, age, and size. For a standard Yorkshire pig, urine flow is estimated to be 0.2–3.2 mL/min with the glomerular filtration rate 1–4.5 mL/min per kg. Blood serum chemistry is similar to that of human blood; for example, creatinine 1.01 ± 0.22 mg/dL, blood urea nitrogen 9 ± 3.2 mg/dL, and calcium 9.6 ± 0.58 mg/ dL (Loeb and Quimby, 1999). A caveat concerning the use of the porcine model is the inflammatory reaction induced by biomaterial implantation. Cormio et al. (1995) have

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shown in 13 piglets that open implantation of various indwelling double-J ureteral stents for 6 weeks resulted in superficial epithelial destruction and ureteral reactive changes in all renal units implanted with stents. The most mild changes occurred with pure silicone stents while the most disruptive were polyurethrane-based stents. The findings differed from their cell culture cytotoxicity studies which showed pure silicone or pure polyurethane stents to be the least cytotoxic while mixtures of silicone or polyurethane did poorly. The authors suggest that the introduction of additives to improve stent handling and biocompatibility have actually worsened the cytotoxic profile, and that possibly the in vivo results were better than expected due to the porcine immune system effectively clearing toxic substances. Olweny et al. (2000) evaluated a prototype lightweight mesh ureteral stent in a porcine model compared with a standard polyurethane 7 F double pigtail stent at 1 and 6 weeks after placement. They noticed that the thinly constructed mesh stent developed a greater weight gain than the 7 F stent at both 1 and 6 weeks. The authors noted no particular encrustation in either group, but a mucinous covering over the mesh stent had developed due to the mucinous metaplasia and mucus production common to indwelling stents placed in the pig ureter. This is distinctly different from the human ureteral reaction as there are no mucus-secreting cells in the human ureter. Desgrandchamps et al. (1995) also noted heavy mucus production in their study of the Wallstent self-expandable metal stent which they cautioned could well prohibit extrapolation of their porcine model data to human results. Olweny et al. (2000) also noted similarly severe ureterectasis and hydronephrosis in both the 7 F polyurethane stent and mesh stent groups due to the high sensitivity of the pig ureter. The work by Sampaio and colleagues has confirmed the pig kidney as an excellent urological model. The pig kidney is very similar to the human kidney in terms of anatomy, physiology, biochemistry, and size (depending on pig size). The animal is hardy and can undergo anesthesia and survive surgeries with relative ease. Any device or material used in the pig can be of the same size used in humans, thus making it much easier to transfer new devices or techniques from the laboratory to clinical application. The caveat is that the pig ureter is more sensitive than the human ureter, and histocompatibility data may not be accurately extrapolated from the porcine model to human results. Additionally, the porcine ureter contains mucinsecreting cells which may confound results regarding patency and flow.

2.3

Conclusion and future trends

While the continued use of in vitro models to test ureteral stents will be of value, only the in vivo model provides a dynamic and robust system for the

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transfer of new devices or materials from the laboratory to the bedside. No single in vivo animal model may take the place of clinical human testing, but a combination of in vivo models to test a single stent may allow the most comprehensive evaluation of the stent in question. Presently, the pig kidney appears to be the single best model for the in vivo testing of ureteral stents with regard to their impact on the kidney; however, tissue reaction to stent material appears to be exaggerated in the porcine model.

2.4

References

atmani, f. and khan, s. r. (1995) Characterization of uronic-acid-rich inhibitor of calcium oxalate crystallization isolated from rat urine. Urol Res, 23, 95–101. atmani, f., opalko, f. j. and khan, s. r. (1996) Association of urinary macromolecules with calcium oxalate crystals induced in vitro in normal human and rat urine. Urol Res, 24, 45–50. banks, r. (1989) Rabbits: models and research applications. USAMRIID Seminar Series, http://netvet.wustl.edu/species/rabbits/rabtmodl.txt. barone, r., pavaux, c., blin, p. and cuq, p. (1973) [In French], Atlas of Rabbit Anatomy, Paris, Masson and Cie, pp. 1–3 and 85–6. bartges, j. w., osborne, c. a., lulich, j. p., kirk, c., allen, t. a. and brown, c. (1999) Methods for evaluating treatment of uroliths. Vet Clin North Am Small Anim Pract, 29, 45–57. bensley, b. (1948) Practical Anatomy of the Rabbit. An Elementary Laboratory TextBook in Mammalian Anatomy, Philadelphia, The Blakiston Company. boyce, w. h., russell, j. m. and webb, r. (1979) Management of the papillae during intrarenal surgery. Trans Am Assoc Genitourin Surg, 71, 76–82. brodsky, s. l., dure-smith, p. and zimskind, p. d. (1977) Gross and radiologic anatomy of the canine kidney. Invest Urol, 14, 356–60. buss, s. l. and bourdeau, j. e. (1984) Calcium balance in laboratory rabbits. Miner Electrolyte Metab, 10, 127–32. cheeke, p. r. and amberg, j. w. (1973) Comparative calcium excretion by rats and rabbits. J Anim Sci, 37, 450–4. cirioni, o., ghiselli, r., minardi, d., orlando, f., mocchegiani, f., silvestri, c., muzzonigro, g., saba, v., scalise, g., balaban, n. and giacometti, a. (2007) RNAIIIinhibiting peptide affects biofilm formation in a rat model of staphylococcal ureteral stent infection. Antimicrob Agents Chemother, 51, 4518–20. cormio, l., talja, m., koivusalo, a., makisalo, h., wolff, h. and ruutu, m. (1995) Biocompatibility of various indwelling double-J stents. J Urol, 153, 494–6. daniel, o. and shackman, r. (1952) The blood supply of the human ureter in relation to ureterocolic anastomosis. Br J Urol, 24, 334–43. denstedt, j. d., reid, g. and sofer, m. (2000) Advances in ureteral stent technology. World J Urol, 18, 237–42. desgrandchamps, f., tuchschmid, y., cochand-priollet, b., therin, m., teillac, p. and le duc, a. (1995) Experimental study of Wallstent self-expandable metal stent in ureteral implantation. J Endourol, 9, 477–81. douglas, g. c. and hossler, f. e. (1995) Vascular anatomy of the rabbit ureter. Anat Rec, 242, 47–56.

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evan, a. p., connors, b. a., lingeman, j. e., blomgren, p. and willis, l. r. (1996) Branching patterns of the renal artery of the pig. Anat Rec, 246, 217–23. evans, h. e., miller, m. e. and christensen, g. c. (1979) Miller’s Anatomy of the Dog, Philadelphia, Saunders. ferber, r., evans, h. and amador, e. (1966) The renal veins of the dog. J Urol, 95, 318–22. finney, r. p. (1978) Experience with new double J ureteral catheter stent. J Urol, 120, 678–81. frodin, l. and engberg, a. (1975) Renal transplantation in the rat. I. Studies concerning the ureteral anastomosis with special reference to the end-to-end technique. Urol Res, 3, 87–90. fuller, p. m. and huelke, d. f. (1973) Kidney vascular supply in the rat, cat and dog. Acta Anat (Basel), 84, 516–22. gibbons, r. p., correa, r. j., jr, cummings, k. b. and mason, j. t. (1976) Experience with indwelling ureteral stent catheters. J Urol, 115, 22–6. gordon, n. (1960) Surgical anatomy of the bladder, prostate gland, and urethra in the male dog. J Am Vet Med Assoc, 136, 215–21. gorman, s. p., garvin, c. p., quigley, f. and jones, d. s. (2003) Design and validation of a dynamic flow model simulating encrustation of biomaterials in the urinary tract. J Pharm Pharmacol, 55, 461–8. harkness, j. e. and wagner, j. e. (1989) The Biology and Medicine of Rabbits and Rodents, Philadelphia, Lea and Febiger. hepperlen, t. w., mardis, h. k. and kammandel, h. (1978) Self-retained internal ureteral stents: a new approach. J Urol, 119, 731–4. hildebrandt, p., sayyad, m., rzany, a., schaldach, m. and seiter, h. (2001) Prevention of surface encrustation of urological implants by coating with inhibitors. Biomaterials, 22, 503–7. hoppe, p. c., laird, c. w. and fox, r. r. (1969) A simple technique for bleeding the rabbit ear vein. Lab Anim Care, 19, 524–5. hossler, f. e. and monson, f. c. (1995) Microvasculature of the rabbit urinary bladder. Anat Rec, 243, 438–48. hossler, f. e. and monson, f. c. (1998) Evidence for a unique elastic sheath surrounding the vesicular arteries of the rabbit urinary bladder – studies of the microvasculature with microscopy and vascular corrosion casting. Anat Rec, 252, 472–6. houston, d. m., moore, a. e., favrin, m. g. and hoff, b. (2004) Canine urolithiasis: a look at over 16 000 urolith submissions to the Canadian Veterinary Urolith Centre from February 1998 to April 2003. Can Vet J, 45, 225–30. jablonski, p., baxter, k., howden, b. o., marshall, v. c., stein-oakley, a. and thomson, n. m. (1995) The effect of ureteric stenting on the function and morphology of long-term rat renal allografts. Aust N Z J Surg, 65, 499–502. kamphues, j. (1991) Calcium metabolism of rabbits as an etiological factor for urolithiasis. J Nutr, 121, S95–6. keane, p. f., bonner, m. c., johnston, s. r., zafar, a. and gorman, s. p. (1994) Characterization of biofilm and encrustation on ureteric stents in vivo. Br J Urol, 73, 687–91. klapproth, h. j. (1959) Distribution of renal arterial circulation in the dog. J Urol, 82, 417–23.

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lee, k. j., johnson, w. d., lang, c. m. and hartshorn, r. d. (1978) Hydronephrosis caused by urinary lithiasis in a New Zealand white rabbit (Oryctolagus cuniculus). Vet Pathol, 15, 676–8. ling, g. v., franti, c. e., ruby, a. l., johnson, d. l. and thurmond, m. (1998) Urolithiasis in dogs. I: Mineral prevalence and interrelations of mineral composition, age, and sex. Am J Vet Res, 59, 624–9. loeb, w. f. and quimby, f. w. (1999) The Clinical Chemistry of Laboratory Animals, Philadelphia, Taylor and Francis. manning, p. j., ringler, d. h. and newcomer, c. e. (1994) The Biology of the Laboratory Rabbit, San Diego, Academic Press. marques-sampaio, b. p., pereira-sampaio, m. a., henry, r. w., favorito, l. a. and sampaio, f. j. (2007) Dog kidney: anatomical relationships between intrarenal arteries and kidney collecting system. Anat Rec (Hoboken), 290, 1017–22. minardi, d., ghiselli, r., cirioni, o., giacometti, a., kamysz, w., orlando, f., silvestri, c., parri, g., kamysz, e., scalise, g., saba, v. and giovanni, m. (2007) The antimicrobial peptide tachyplesin III coated alone and in combination with intraperitoneal piperacillin-tazobactam prevents ureteral stent Pseudomonas infection in a rat subcutaneous pouch model. Peptides, 28, 2293–8. mitruka, b. m. and rawnsley, h. m. (1977) Clinical Biochemical and Hematological Reference Values in Normal Experimental Animals, New York, Masson Publ. USA. miyake, o., yoshioka, t., yoshimura, k., honda, m., yamaguchi, s., koide, t. and okuyama, a. (1998) Expression of Tamm-Horsfall protein in stone-forming rat models. Br J Urol, 81, 14–9. national research council (us), committee on the handbook of biological data and spector, w. s. (1956) Handbook of Biological Data, Philadelphia, Saunders. olweny, e. o., portis, a. j., sundaram, c. p., afane, j. s., humphrey, p. a., ewers, r., mcdougall, e. m. and clayman, r. v. (2000) Evaluation of a chronic indwelling prototype mesh ureteral stent in a porcine model. Urology, 56, 857–62. orlando, f., ghiselli, r., cirioni, o., minardi, d., tomasinsig, l., mocchegiani, f., silvestri, c., skerlavaj, b., riva, a., muzzonigro, g., saba, v., scalise, g., zanetti, m. and giacometti, a. (2008) BMAP-28 improves the efficacy of vancomycin in rat models of gram-positive cocci ureteral stent infection. Peptides, 29, 1118–23. pfeiffer, e. w. (1968) Comparative anatomical observations of the mammalian renal pelvis and medulla. J Anat, 102, 321–31. sampaio, f. j. and aragao, a. h. (1990) Anatomical relationship between the intrarenal arteries and the kidney collecting system. J Urol, 143, 679–81. sampaio, f. j. and favorito, l. a. (1993) Ureteropelvic junction stenosis: vascular anatomical background for endopyelotomy. J Urol, 150, 1787–91. sampaio, f. j., pereira-sampaio, m. a. and favorito, l. a. (1998) The pig kidney as an endourologic model: anatomic contribution. J Endourol, 12, 45–50. seaman, r. and bartges, j. (2001) Canine struvite urolithiasis. Compend Cont Educ Pract Vet, 23, 407–20. sheehan, h. l. and davis, j. c. (1959) Anatomy of the pelvis in the rabbit kidney. J Anat, 93, 499–502. smith, m. j. and boyce, w. h. (1968) Anatrophic nephrotomy and plastic calyrhaphy. J Urol, 99, 521–7.

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tunney, m. m., bonner, m. c., keane, p. f. and gorman, s. p. (1996) Development of a model for assessment of biomaterial encrustation in the upper urinary tract. Biomaterials, 17, 1025–9. wolf, j. s., jr, humphrey, p. a., rayala, h. j., gardner, s. m., mackey, r. b. and clayman, r. v. (1996) Comparative ureteral microanatomy. J Endourol, 10, 527–31. yasuda, h., koga, t. and fukuoka, t. (1999) In vitro and in vivo models of bacterial biofilms. Methods Enzymol, 310, 577–95. zimskind, p. d., fetter, t. r. and wilkerson, j. l. (1967) Clinical use of long-term indwelling silicone rubber ureteral splints inserted cystoscopically. J Urol, 97, 840–4.

3 Models for the assessment of biofilm and encrustation formation on urological materials B. F. G I L M O R E, D. S. J O N E S and S. P. G O R M A N, Queen’s University Belfast, Northern Ireland; and H. C E R I, University of Calgary, Canada

Abstract: Medical devices of the urinary tract are in commonplace usage in modern urology and are likely to remain the cornerstone of many surgical interventions in the field. This is despite the fact that their effective use is still drastically hindered by the formation of biofilm and encrustation which can cause obstruction and blockage of the device, as well as significant morbidity in the patient. This chapter examines the in vitro encrustation models available for evaluation and preliminary assessment of new biomaterials, coatings and drug-eluting devices for use in the urinary tract, aimed at resisting surface encrustation and microbial biofilm formation. Key words: encrustation, in vitro models, Proteus mirabilis, struvite, hydroxyapatite, ureteral stent, urethral catheter.

3.1

Introduction

Ureteral stents and urethral catheters are employed to facilitate urine flow and drainage from the upper and lower urinary tracts, respectively.1 Ureteral stents, which have become a cornerstone of contemporary urological practice, are indicated for the prevention or relief of ureteral obstruction and facilitate the flow of urine from the kidney to the bladder.1–4 Typical indications for stent placement include obstructing ureteral calculi, ureteral strictures, carcinomas, retroperitoneal fibrosis, trauma, congenital anomalies, obstructive uropathy and facilitation of reconstructive surgery and transplantation.2,5 Indwelling urethral catheters are utilized in the management of an array of both chronic and acute conditions, generally indicated where mobility is compromised and/or there is damage to the physiology of the bladder or urethra. Some indications for urinary catheterization include urological surgery, urinary output monitoring, acute urinary retention, bladder irrigation, incontinence, palliative care for terminally ill or severely impaired incontinent patients and uncorrected prostatic obstruction.6 59

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The use of prosthetic medical devices to support the normal physiological functions of the urinary tract is by no means a modern concept, with the first historical reference to their use dating back over two millennia to the Greek physiologist Erasistratus of Ceos (310–350 BC), who is credited with the invention of a urinary tract catheter for the treatment of urinary retention.7 The first recorded case of ureteral stenting was performed by Gustav Simon in the nineteenth century.5,7 The era of modern long-term indwelling ureteral stents began in the late 1960s, thanks to the pioneering work of Zimskind and colleagues.8 In more recent times, instrumentation of the urinary tract has come to represent the most common and widespread use of implantable medical devices, which have become an integral aspect of urological practice worldwide. Unfortunately, their use is associated with a high incidence of complications common to many implantable, indwelling medical devices. These are primarily bacterial infection and colonization, leading to microbial biofilm formation and encrustation, a process whereby inorganic salts are deposited on the surface or lumen of the medical device. The range of devices routinely used within the urinary tract is extensive, with urethral catheters and ureteral stents still accounting for the overwhelming majority of such devices.6 It is estimated that some 11% of hospitalized patients in Europe host a urinary catheter at some stage during their stay in hospital, with approximately 5 million patients in acute care hospitals catheterized annually.9 Interestingly, levels of catheterization in the USA are significantly higher than in Europe, with estimates ranging from 16 to 25% of hospitalized patients.10,11 Somewhere in the region of 5% of residents in nursing homes are catheterized long-term.12 Long-term catheterization (LTC), more common in those patients requiring residential care in nursing homes, is generally reserved for those patients for whom all other measures have failed because of the high risk of complications.13 The prosthetic medical devices of the urinary tract employed in modern urology are fabricated from a range of polymeric materials including latex, silicone, silicone co-polymers and polyurethane.3,14 Despite this diversity of materials currently available – each offering unique chemical, physical and mechanical property sets – no perfectly biocompatible material that is not subject to modification by its environment has been defined.15 For example, Tunney et al. report that no single material – including polyurethane, silicone, composite biomaterials and hydrogel-coated material – was more effective than any other in preventing microbial colonization.16 Therefore, despite improved design and technological innovation in biomaterial design, the incidence of encrustation and attendant complications remains high. Following implantation, urological devices rapidly succumb to surface modification, initially by an adsorbed conditioning film composed of

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inorganic ions, as well as host- and bacterial-derived extracellular matrix proteins, glycoproteins and exopolymers, all of which can mask the underlying characteristics of the device,17,18 and facilitate subsequent microbial biofilm formation. Currently a biofilm is defined, by Donlan and Costerton as ‘a microbially derived sessile community characterized by cells that are irreversibly attached to a substratum or interface or to each other, are embedded in a matrix of extracellular polymeric substances that they have produced and exhibit altered phenotype with respect to growth rate and gene transcription’.19 Bacteria present in this biofilm are able to initiate the infection process and the biofilm itself can act as the nidus for encrustation and a reservoir for infection.20,21 Furthermore, a general characteristic of biofilm communities is their tendency to exhibit significant tolerance to antimicrobial challenge compared with planktonic bacteria of the same species, with bactericidal/eradication concentrations of antimicrobials being as much as 1000 times higher for biofilms than for planktonic bacteria.22 The inaccessibility and inherent resistance of the device-related biofilm to antimicrobials constitutes the fundamental difficulty in dealing effectively with such device-associated infections. Management of biofilm and encrustation on medical devices is therefore problematic and, in a majority of cases, necessitates complete removal of the device. Although catheters thus fouled are able to be removed and replaced with relative ease and low cost, removal of encrusted stents may require major surgery and associated trauma for the patient and cost implications for providers of healthcare.16 The use of medical devices of the urinary tract, while often providing simple and cost-effective solutions to a range of chronic and acute aetiologies of the urinary tract, has however been demonstrated to carry with it significant impacts on quality of life, increases in patient morbidity23,24 and mortality.25 Regrettably, up to 80% of ureteral stent patients report reduced health-related quality of life scores.23 Clinical need dictates that biomaterials used in the manufacture of urinary tract medical devices should resist bacterial colonization and encrustation and retain their functional characteristics for the intended lifetime of the device. To this end, advances continue to be made in the areas of materials science, drug development, microbiology and clinical medicine. This has given rise to a new generation of devices designed to counter these problems and thus bring about improved patient outcomes by enhancing the patients’ quality of life and reducing morbidity and mortality rates in patients requiring such medical device placement. Reliable and robust in vitro models are therefore required for evaluating the suitability and, preliminary assessment of new biomaterials, coatings and drugeluting devices aimed at resisting surface encrustation and microbial biofilm formation.

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3.2

Development of urinary encrustation

Encrustation of urinary devices with mineralized deposits remains one of their primary complications with estimates of the overall incidence as high as 58%.10,26 There are many patient-dependent variables affecting the formation of encrustations with one of the key factors determining propensity towards encrustation being the length of time that the device remains within the urinary tract. Studies suggest that some 50% of patients undergoing catheterization for more than 28 days develop recurrent encrustation or blockage11 while as many as 75% of stents are encrusted after 24 weeks.27 A recent report noted that patients catheterized for 76 or more days were 3 times more likely to die, be hospitalized or require antibiotics than matched control patients.25 El-Faqih and colleagues28 reported that of 141 retrieved stents subjected to luminal patency evaluation, 46.8% showed impassable blockage with encrusting deposits. Furthermore, they reported that stent encrustation increased from 6% in stents retrieved before 6 weeks to 76% in stents indwelling for more than 12 weeks. Additional studies indicated that between 44 and 69% of stents recovered from patients had bacterial colonies present, with some 29% of these patients having been diagnosed with bacteriuria requiring treatment.29,30 Highest rates of colonization (75–100%) not prevented by oral antibiotic therapy, were observed in patients with stents indwelling for longer than 3 months.29,30 These deposits frequently obstruct the lumen of devices leading to urinary retention, painful distension of the bladder or more severe complications such as urolithiasis, pyelonephritis, hepatic encephalopathy, septicemia and shock.31 In addition, the abrasive nature of encrusted urinary devices may cause trauma on device removal and lead to permanent damage to the uroepithilium.1 Although encrustation can occur in the absence of infection (metabolic encrustation), it is primarily associated with microbial biofilm formation and, as such, figures for its financial impact alone are not available, although inferences can be drawn from the costs of urethral catheter-associated infections. Current estimates of the combined cost of managing encrustation and urinary tract infection associated with indwelling urinary devices amounts to somewhere in the region of £1 billion annually in Western Europe with a similar, if not higher, burden being placed upon the health system in the USA.12 The Gram-negative organisms most commonly implicated as the primary pathogenic species are Escherichia coli, Proteus mirabilis, Pseudomonas aeruginosa and Klebsiella pneumoniae. Despite this, Gram-positive strains such as Staphylococcus epidermidis and Enterococcus faecalis are also commonly implicated.32 Infection of the urinary tract is dependent on the time that a device remains indwelling. For example, 10–50% of catheterized

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patients have developed clinically significant bacteruria within 7 days and by 28 days virtually all patients are colonized with at least one of the aforementioned pathogenic species.32 Encrustations associated with an infectious origin are primarily composed of magnesium ammonium phosphate (struvite) [MgNH4PO46H2O] or calcium phosphate (hydroxyapatite) [Ca10(PO4)6(OH)2] and may affect both indwelling urinary catheters and stents. Numerous strains of bacteria are known to produce the enzyme urease during pathogenesis; however, the key pathogen in the development of encrustation is Proteus mirabilis.33 This ability to produce encrustation in the urinary tract is due to several key virulence factors. P. mirabilis hydrolyses urine some 6–10 times faster than these other microorganisms34 with its optimal activity at pH 7 reduced by approximately 50% at pH 9.35 Swarmer cells, induced by conditions that inhibit flagellar rotation such as contact with a material surface as in biofilm formation,36 are associated with a 30-fold increase in the production of urease and the production of polysaccharide ‘slime’ which has been demonstrated to promote struvite crystal growth.20,37 The composition of the concretion arising from encrustation of infectious origin is similar to urinary stones38 with the same bacterial strains being implicated.39 The process of encrustation has been well documented33 and can be broken down into several fundamental steps. Colonization of the biomaterial surface with urease-producing bacteria causes alkalinization of the urine and the biofilm matrix, thus lowering the solubility of struvite and hydroxyapatite in the urine and the immediate milieu of the biofilm. The urease-mediated hydrolysis of urea liberates ammonia and carbon dioxide: H2N H2N

O

+

H2O

Urease

2NH3 + CO2

Under neutral or slightly acidic conditions (i.e. normal urine pH) the ammonia becomes protonated: NH3 +

H2O

NH4+ + OH-

Carbon dioxide reacts with water to produce carbonic acid, which may dissociate depending on urinary pH: CO2 + H2O

H2CO3

HCO3- + H+

CO32- + 2H+

Hydroxide formation from the reaction of ammonia and water, despite the presence of weakly acidic carbonic acid, results in an overall elevation of urine pH and the subsequent precipitation of struvite and hydroxyapatite from urine which are then deposited on the biomaterial surface. Figure 3.1

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(b)

3.1 Scanning electron microscopy image of hydroxyapatite and struvite (coffin-shaped crystals) encrustation on the surface of a Foley catheter. Panel (b) is reproduced from reference 6 with permission from Future Drugs Ltd.

shows scanning electron microscopy images of hydroxyapatite and struvite encrustation on the surface of a Foley catheter. The most common ureaseproducing uropathogens are Proteus mirabilis, P. vulgaris and Pseudomonas spp.40 Furthermore, Escherichia coli has also been shown to increase the rate of urease-dependent encrustation in an in vitro model.41

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Assessment of biomaterial encrustation – in vitro models

In vitro encrustation models permit the assessment of urinary tract biomaterials and as such represent a cost-effective and reliable method for undertaking preliminary evaluation of novel approaches to the development of materials that resist these deleterious processes. Since effective strategies for the prevention of bacterial biofilm formation and encrustation on medical devices for use in the urinary tract would represent a major breakthrough in patient care and quality of life, such models are valuable tools in the design and validation process. These models have also demonstrated utility in the interpretation of clinical data and observations; for example, in the use of prophylactic antimicrobial cover.42 The experimental models of encrustation are based on pooled human urine or artificial synthetic urines. Pooled human urine suffers primarily from patient-dependent variability of composition (inter- and intra-subject variability); however, it has advantages in that the system may be more representative of the in vivo conditions within the urinary tract since urine is a complex solution, the properties of which depend on the interaction of components of the solution rather than being simply an aggregate of the properties of the individual components.1,43,44 In contrast, artificial urine composition can be tightly controlled; batch-tobatch experimental variation can be eliminated and supply is not limited, allowing large numbers of samples to be tested at a given time. The majority of studies on in vitro encrustation models that do not rely on pooled human urine employ artificial synthetic urines, essentially either as described by Griffith et al.45 or a simplified, urease-containing formula devised by Cox et al.46 The compositions of these synthetic artificial urines are given in Tables 3.1 and 3.2, respectively. Stickler and colleagues

Table 3.1 Composition of artificial urine devised by Griffith et al.45 Component

Concentration (g/l)

Calcium chloride dihydrate Magnesium chloride hexahydrate Sodium chloride Sodium sulphate Sodium citrate dihydrate Sodium oxalate Potassium dihydrogen orthophosphate Potassium chloride Ammonium chloride Urea Creatinine

0.65 0.65 4.60 2.30 0.65 0.02 2.80 1.60 1.00 25.0 1.10

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Table 3.2 Composition of artificial urine devised by Cox et al.46 Solution

Component

Concentration (g/l)

Solution 1

Potassium dihydrogen orthophosphate Magnesium chloride hexahydrate Urea Calcium chloride hexahydrate Chicken ovalbumin Jack bean urease type IX

7.62

Solution 2 Urease

3.64 16.00 5.30 20.00 0.04

demonstrated the importance of urine conditioning film, and thus modification of the biomaterial surface after implantation, on the formation of a range of microbial biofilms,42 findings confirmed by Santin and co-workers.18 Many studies report the use of sterile pooled human urine, infected human urine,47 artificial synthetic urines48,49 or urine to which urease has been added to mimic the effect of the presence of urease-splitting (i.e. ureaseproducing) microorganisms.46,50 Numerous models have been described that simulate encrustation in the upper and lower urinary tracts and acknowledge the site-specific flow characteristics of urine in the urinary tract. Urine presents in the urinary tract either as a relatively static body, as in the bladder, or as a flowing fluid in the ureters and urethra. Models with urine movement, or dynamic models, as well as static models have been described in the literature. These are used to study encrustation on ureteral stents and urethral catheters, respectively.

3.4

Dynamic flow-through models

One of the first dynamic flow-through in vitro models for the study of encrustation, specifically calcium oxalate deposition, was described by Finlayson and Dubois.51 Their model, a modification of an earlier experimental design,52 permitted the study of calcium oxalate encrustation on, and dissolution from, glass rods. The reaction chamber, a double-walled glass vessel with an inflow port at the top and an outflow spigot from the inner vessel, was kept at a constant temperature by water pumped through the outer jacket from a heated circulator. The contents of the reaction vessel were mixed with a magnetic stirrer and, as such, the effect of stirring speed on deposition and crystal growth could be evaluated. Rods to be encrusted were suspended in a 550 ml-capacity continuous flow vessel containing 275 ml of 0.5 mM K2C2O4 and 275 ml of 5.0 mM CaCl2, which had been stirred in the vessel for 24 h prior to insertion of the glass rods. Solutions of Tris-buffered saline plus K2C2O4 (3.1 mM) and CaCl2 (20 mM), pH 7.5,

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Solution input lines

To constant temperature

Outflow

From constant temperature

3.2 Finlayson and Dubois continuous flow vessel for the study of encrustation studies. Adapted from Finlayson and Dubois.51

are simultaneously pumped into the system, at an inflow rate of 300 ml/h. The authors report experimental data over a 40 h period. A schematic diagram of the Finlayson and Dubois model is shown in Fig. 3.2.51 The reproducibility of this method was found to be sufficiently robust to allow preliminary examination of temperature, stirring speed and flow rate. This system is widely regarded as the basic apparatus for simple encrustation studies. Numerous modifications to this original design have been devised to facilitate the study of urinary encrustation and to circumvent a number of inherent problems of using the original system (especially in the presence of urease-producing microorganisms); these modifications are described in the literature.45–47,49 Gleeson and co-workers47 proposed a modification of the Finlayson and Dubois model that consisted of a 700 ml glass chamber with inflow and outflow ports, with the system maintained at 37 °C on a stirrer hot plate. This adapted model accommodated a daily inflow rate of 500 ml and had a turnover rate of 4 days. Pre-weighed biomaterial discs (6 × 0.5 mm) were suspended from silicone rods in the glass vessel which was sterilized by autoclaving. Encrustation studies are initiated by filling the glass vessel to capacity and maintaining the daily flow rate. All biomaterial discs were dried and weighed to determine mass of deposition before further analysis by optical and radiograph diffraction crystallography. This model permitted the comparative evaluation of the encrustating behaviour of a range of biomaterials in both sterile and infected (urease-producing Proteus

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vulgaris) pooled urine. The authors report that silicone–carbon composite encrusted to a lesser extent than the other materials tested and also concluded that the degree of encrustation of the biomaterials under examination depended on the biomaterial composition, length of urinary exposure, presence of infection and solute content of the urine. This dynamic flowthrough system permitted examination of encrustation over a period of 2 weeks, a significant improvement on the previous model. The slow turnover rate of this model, and its ability to assess only small numbers of biomaterials at a time have been identified as this model’s key disadvantages.4 A further improvization of the model is described by Sarangapani and co-workers,49 in a study that highlighted a number of problems encountered in the routine use of the Finlayson and Dubois model with Proteus mirabilis-infected synthetic urine, as well as defining a number of adaptations to circumvent such complications. The reactor consisted of a 500 ml glass beaker reaction vessel, with an inflow channel situated on the vessel lid and an outflow channel on the side of the reactor. The reactor lid was designed with a central opening to accommodate the inlet tube and six evenly spaced silicon septum-sealed openings through which pre-weighed sample biomaterials (catheter sections) were suspended in the reaction vessel. Synthetic urine, as described by Griffith et al.,45 also containing 5% tryptic soy broth, was supplied to the vessel from a reservoir and displaced urine collected in a waste container. Bacterial air vent filters were also added to the modified system to prevent contamination of the sterile urine in the reservoir, the reaction vessel and the waste container. The reaction chamber was filled with 495 ml sterile synthetic urine, and perfused with sterile synthetic urine at a rate of 50 ml/h until the inlet lines were purged of air. Reactor chambers for the comparative assessment of infectionrelated encrustation were then inoculated with 5 ml P. mirabilis culture (1 × 105 colony forming units (cfu)/ml). Perfusion was continued with sterile synthetic urine containing 5% tryptic soy broth. Encrustation of catheter materials in this study was assessed by increase in mass after the experimental test period, with results indicating that, over a 7 day period latex and hydrogel-coated catheters encrusted to the same extent. In essence, the original model required modifications to permit extended experimental periods, since using a straight-inlet flow tube resulted in severe encrustation and blockage of the inlet tube lumen by 48–72 h, thus preventing free drainage of urine from the reservoir to the reaction vessel. Furthermore, bacteria were found to ascend the feed tube from the reaction vessel to the pump causing further encrustation. The inlet apparatus of the device was modified to incorporate an ‘anticlimb’ top joint, which consisted of an S-shaped tube housed in a large glass bulb, in order to break liquid contact between the reservoir and the infected urine in the reaction vessel. A disadvantage of this experimental model is the relative complexity of the design and, similar

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to the Gleeson model described above, the lack of capacity for multiple or replicate samples of material. The simple physical model for the study of biofilm/encrustation formation on urethral catheters described by Stickler and co-workers42,53 represents a significant development in experimental design whereby the hydrodynamic environment of the catheterized bladder is effectively mimicked. Essentially, the model consists of a 200 ml glass fermentation flask maintained at 37 °C by a water jacket. After assembly, the model is sterilized by autoclaving and a catheter (ⱅ14) is aseptically inserted into the flask via a section of silicone tubing (the in vitro ‘urethra’) attached to the glass outlet in the base of the flask. The catheter retention balloon is inflated by the addition of 10 ml of water, thus securing the catheter in place and sealing the outlet from the vessel. After attaching the catheter to a drainage bag, the sterile urine enters the bladder from the inlet port via a peristaltic pump, where a residual volume of 30 ml accumulates in the ‘bladder’ beneath the level of the eyelet. Urine overflow drains through the catheter eyelet and collects in the reservoir bag, under gravity. The model of the catheterized bladder is shown in Fig. 3.3. Pooled human or artificial urine may be used as a growth medium for organisms under test. The artificial urine used in the model is a modification of that described in reference 45.

Urine Sampling port

Water, 37°C Glass outlet tube Silicone tubing

Tube to drainage bag

3.3 Simple physical model of the catheterized bladder. Adapted from Stickler et al.42,53

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The inoculum for the vessel (10 ml of early exponential phase culture of the test organism in urine) is added to 20 ml of the residual volume in the vessel and allowed to incubate for 1 h before commencing supply of urine to the flask at a rate of 0.5–1.0 ml/min. This arrangement has been used to study the ability of a wide range of organisms to produce catheter biofilms and concretion. Bladder populations of 107–108 cfu/ml can be maintained in this way for up to 10 days. The catheter is removed from the bladder model at the end of the experimental period by deflation of the retention balloon and lumenal encrustation may be examined using various analytical and visual assessment techniques. The model has provided much useful information in relation to the formation, physiology and control of catheter biofilms. For example: in the identification of species of bacteria that are capable of producing crystalline biofilms and hence catheter blockage;33,54 the study of crystal formation and structure of Proteus mirabilis catheter-encrustating biofilms;55 the effect of pre-treatment with antimicrobials and the importance of the urine conditioning film;55,56 assessing which commercially available indwelling urethral catheters are capable of resisting encrustation by P. mirabilis biofilm57 to address critical questions in Foley catheter design, for example lumenal surface characteristics, which make these devices so vulnerable to encrustation;58 and evaluation of novel approaches to encrustation-resistant catheter design.59 Choong and colleagues44 describe, in a further adaptation, a versatile, dynamic model to quantify encrustation on ureteric stents, urethral catheters and polymers for urological uses. The model consists of a ‘bladder’ reservoir, supplied with (single-source, antibiotic-treated) human urine from a central urine reservoir at a constant flow rate of 0.5 ml/min through silicone tubing via the inlet of the model. The outlet of the ‘bladder’ reservoir is adapted to house either a urethral catheter or a siphon which controls the accumulation of urine in the vessel such that the bladder empties when a defined volume (dictated by the height of siphon) has been reached. In this way voiding of the urine chamber is designed to mimic human micturation. A further modification, a side-arm, which permits the attachment of a ‘ureteric’ section to an upper ‘kidney’ vessel permits adaptation of the model to accommodate a ureteral stent. The entire device is housed within a carbon dioxide incubator to permit maintenance of temperature, pH, gaseous composition and humidity. Materials/devices are exposed to urine continuously for 5 days. Samples retrieved from the model are analysed by atomic absorption spectroscopy and results expressed as an encrustation index, defined as the ratio of encrustation of the test and that of a reference material. This model has been used to assess the effects on encrustation of calcium leaching from glass vessels and the use of single-source and pooled human urine, and to quantify directly the encrustation on commercially available ureteric stents.

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3.5

71

Batch flow or ‘static’ models

The models described below are referred to in the literature as ‘static’ models. However, these models employ a batch culture, where turnover of fluid is not constant; shear force is created in these models to emulate the flow of urine as would be seen in a catheter system, hence these are not truly static models. Tunney et al.50 described a batch artificial urine model for the study of urinary encrustation in the upper urinary tract. The reaction vessel, as shown in Fig. 3.4, consists of a Perspex tank, with a loose-fitting lid. The artificial urine employed in this model is based on that described by Cox et al.46 but with substantially reduced albumin content. Solutions are added separately to the reaction vessel to prevent acidic precipitation of brushite. The artificial urine (5.16 l) is stirred by means of Teflon-coated metal stirring bars to induce shear force on the catheters, as would be seen in a flow system. A plastic grid is positioned 80 mm above the tank floor from which biomaterial sections (length, 50 mm) are suspended in the artificial urine by means of colour-coded, plastic-coated paper clips. This allows several biomaterials to be evaluated concurrently. An aperture in the plastic grid permits daily solution exchange – 1 l is removed daily using a siphon pump and replaced with an equivalent volume of pre-heated (37 °C) artificial urine. Twice weekly, 160 ml is removed and replaced with an identical volume of a third solution containing urease. Sections are removed at defined test periods (i.e. 2, 6 and 14 weeks), rinsed with deionized water, and the type and quantity of encrustation determined by infrared spectroscopy, X-ray diffraction spectroscopy, energy dispersive X-ray analysis and atomic absorption spectroscopy. This model permits large numbers of

5.16 litres artificial urine 37°C, 5% CO2

3.4 The static artificial urine model for assessment of biomaterial encrustation. Adapted from Tunney et al.50

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biomaterials to be evaluated simultaneously and over extended time periods.50 Results from this study indicated that the type of encrustation produced on polyurethane stents in vitro was composed primarily of struvite and hydroxyapatite, and was similar to that produced on stents in vivo.10 In a similar model Jones et al.60 describe a batch, artificial urine bladder encrustation model to study the effects of the components of artificial urine and urease inhibitors on encrustation produced on commercially available stent materials. In essence, the model consisted of a 700 ml-capacity plastic reaction vessel with a firmly attached lid. A schematic diagram of the bladder encrustation model is shown in Fig. 3.5. Sections (2.5 cm) of the ureteral stents were dissected, heat-sealed at both ends and suspended into the artificial urine using Microlance 3 needles (60 mm length) which were firmly attached within the vessel. An aperture on the reaction vessel lid permitted daily replacement of (120 ml) artificial urine. The vessel(s) were placed in an orbital shaker and their contents were exposed to a physiological temperature of 37.0 ± 0.1 °C and a rotation speed of 100 rotations/min. The composition of the artificial urine used in this study was adapted from Tunney et al.50 After 7 days, the masses of calcium and magnesium encrustation on the surface of the stents were quantified by atomic absorption spectroscopy. The model was used to examine the effects of individually altering the concentrations of calcium chloride hexahydrate (0–0.53% w/v) and chicken ovalbumin (0–0.2% w/v) within artificial urine, the

3.5 Static encrustation vessel containing five stent sections. Figure reproduced from reference 60 with permission from Wiley Periodicals Inc.

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concentration of magnesium chloride hexahydrate (0–0.36% w/v) and the presence or absence of urease on the mass of calcium and magnesium encrustation on the selected biomaterials. The pH values of the various artificial urine solutions were monitored potentiometrically. Finally, in order to examine the effects of various urease inhibitors on encrustation on the various biomaterials, methylurea, ethylurea or acetohydroxamic acid (each 10 mM) were incorporated into the standard artificial urine and the mass of calcium and magnesium encrustation determined following immersion within the encrustation model for 1 week. This model exhibited high reproducibility, attributed to improved stirring conditions (since the entire vessel is placed on a gyrorotary platform), and also demonstrated the potential for reducing urinary encrustation by incorporation of agents that modify urease activity, such as acetohydroxamic acid, into biomaterials.

3.6

Dynamic continuous flow models

Continuous flow models are more representative of portions of the urinary tract, such as the ureters, where urine flow is continuous. Such models are useful in the evaluation of the ability of a biomaterial to resist encrustation and the comparative ability of ureteral stents to resist intraluminal blockage. Tunney and colleagues61 describe a continuous flow artificial urine model employing a modified Robbins device (MRD). The Robbins device was developed at the University of Calgary62 to study fouling biofilm formation in industrial water systems before being modified to study biofilms of medical importance, and has been used widely in the study of biofilm formation and control. The MRD is an acrylic block, 42 cm long with a 2 × 10 mm lumen. Twenty-five evenly spaced sampling ports are designed to allow the attachment of biomaterial discs to sampling plugs housed in each of the ports. Biomaterial discs are attached to the plugs using a rubber backing sheet and lie flush with the inner lumen so as not to disturb flow characteristics. Artificial urine, as described in reference 50, is pumped at a rate of 0.7 ml/min (simulating normal urine flow) from a reservoir equilibrated at 37 °C/5% CO2 through the MRD and back to the reservoir. A schematic diagram of the recirculating MRD encrustation model is shown in Fig. 3.6. This model has been used to study encrustation on silicone and polyurethane, revealing significantly less deposition of hydroxyapatite and struvite on silicone than on polyurethane discs.1 In a further modification of this experimental model, a simulated urine flow model is also described.16 In this model, urine is pumped from the reservoir through 60 cm lengths of ureteral stent sections, rather than the MRD. Stent sections are checked daily to ensure patency and urine flow. If urine flow through the section ceases owing to encrustation and blockage of the lumen, the section of biomaterial is removed, dried and the total mass

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Biomaterials and tissue engineering in urology

Urine flow rate 0.7 ml/min Artificial urine reservoir 37°C, 5% CO2

Modified Robbins device Peristaltic pump

3.6 Dynamic recirculating (modified Robbins device) encrustation model. Adapted from Tunney et al.61

of encrustation determined. This model has been used to assess the encrustation of materials previously evaluated in the batch model described above.50 The dynamic encrustation model described by Gorman and co-workers4 consists of a purpose-designed glass reaction vessel linked to a peristaltic pump to circulate artificial urine through the vessel which contains biomaterial samples attached to stainless steel mandrels, as shown in Fig. 3.7(a) and (b). Artificial urine was pumped from a 5.0 l reservoir of artificial urine through silicone tubing (inner diameter 7.7 mm, wall thickness 1.3 mm). The vessels were filled initially with urine using a plastic syringe. Gate clamps were used to temporarily seal the tubing leading from the vessels to prevent backflow of urine as the vessels were clamped into position. The inlet tube, pump manifold tubing and tubing leading to the vessels were filled with urine from the reservoir by turning on the pump. The gate clamps were then removed. A flow rate of 10 ml/min was necessary to maintain the urine level in the reaction vessel. Several reaction vessels may be used in parallel. Each vessel had a separate inflow and outflow of urine and outflow urine was pumped back into the reservoir. Biomaterial sections to be tested were secured equidistant along each mandrel using adhesive. Up to three sections of biomaterial (3 cm) per mandrel and three mandrels per vessel may be tested. The mandrels were secured into rubber stoppers at each end of the vessel. The entire system was maintained at 37 °C to mimic physiological temperature. Artificial urine (1.0 l) was replaced in the reservoir daily and the operation of the model was checked to ensure reproducible flow from the outlet tube over a 5 min period. Any crystalline material deposited in the tubing was also cleared daily to prevent blockages. The reproducibility of the model was validated to ensure that each biomaterial position resulted in

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(a) Inflow of urine (10 ml/min) Glass reaction vessel filled with artificial urine Mandrels and material sections

Outflow of urine (10 ml/min) (b)

Pump

Artificial urine reservoir

3.7 Diagrammatic representation of (a) the reaction vessel and (b) the urine pumping system in the dynamic encrustation model. Figure reproduced from reference 4 with permission from Pharmaceutical Press.

statistically similar levels of encrustation. The performance of the dynamic model was validated initially by comparative masses of encrustation (hydroxyapatite and struvite) on biomaterials that are commonly employed in urinary tract. After 2 weeks of urine flow, calcium and magnesium encrustation on each section of material were determined. Further validation of the model examined encrustation deposited on each of the biomaterials (polyurethane, silicone) in the dynamic model compared with the encrustation arising in the batch encrustation model described in reference 50. The authors conclude that this model provides a reproducible method of mimicking encrustation formation on biomaterials in portions of the urinary

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tract experiencing dynamic flow conditions. This model is particularly applicable to the evaluation of drug-eluting or biodegradable/surface-shedding biomaterials intended for use in the urinary tract.

3.7

The MBEC-BESTTM assay

One of the newest technologies available to study biofilm formation and encrustation of catheters and stents is the BESTTM assay, a patented modification of the MBECTM (minimum biofilm eradication concentration) assay technology63,64 available through Innovotech Inc. (www.innovotech.ca). Essentially the BESTTM assay makes use of the matrix capabilities of the MBECTM system; however, the pins of the MBECTM system are replaced as a surface for biofilm formation/encrustation by any surface of interest for study, and in this case by the device in question – a catheter or stent surface. The ability to adapt the BESTTM assay to almost any surface shape or type makes this one of the most flexible biofilm/adhesion/encrustation assays available. The BESTTM assay consists of a lid that is configured to support the material in question, while the bottom is made up of a welled plate into which the surface in question will sit. Figure 3.8(a) demonstrates the lid of

(a)

(b) Mounting pins Test stent

BEST lid

Artificial urine and bacteria

3.8 (a) BESTTM assay lid configured for the testing of catheters or stents; (b) schematic diagram of the BESTTM assay plate.

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the BESTTM assay configured for the testing of catheters or stents. The matrix and high-throughput capabilities of the BESTTM assay allows one to evaluate the effect of a number of parameters on encrustation. One can evaluate encrustation in the presence of a number of different bacterial species or in bacteria-free systems. Growth parameters such as pH, oxalate levels, artificial urine (as compared with bacterial growth media), presence of host serum and tissue factors can all be easily measured in any set of combinations. The experimental set-up is very simple; catheter sections are attached to the BESTTM lid via surface pins present on the lid of the device that can accommodate catheters or stents of various diameters. The mechanism of placement of the catheters and stents onto the BESTTM lid allows contact only with the coated surfaces of the device, thereby preventing contact with the cut ends of the catheter/stent and thus avoiding complications that may arise when testing surface coatings for their efficacy in preventing biofilm formation and encrustation. The lid bearing the devices to be tested is then placed over the appropriate commercially available 12-, 24- or 36-well plate. Each well contains the test solutions as required (as shown in Fig. 3.8(b) where artificial urine is the test medium). The BESTTM device is then placed on a gyratory shaker at a desired speed, in an incubator at the desired temperature for the time required. Samples can be taken at time periods by simply removing the lid, collecting catheter sections and carrying out standard chemical analysis of encrustation products. Biofilm formation can be determined by placing the lid over a collection plate and sonicating the bacteria into recovery media for enumeration. Encrustation can be followed for days to weeks by replenishing the volume of the growth medium or by replacing the well plate with one containing fresh incubation media.

3.8

Conclusions

Medical devices of the urinary tract remain in commonplace usage in modern urology and are likely to remain the cornerstone of many surgical interventions in the field. This is despite the fact that their effective use is still drastically hindered by the formation of biofilm and encrustation which can cause obstruction and blockage of the device, as well as significant morbidity in the patient. Clinical need is driving, and will continue to drive, the evolution of new-generation medical devices of the urinary tract. In vitro models of urinary encrustation of biomaterials, as described in this chapter, remain the best option for the assessment of new approaches towards the development of truly biomimetic and biocompatible biomaterials that will effectively resist infection and encrustation, and permit permanent or long-term placement free from attendant complications. Various new approaches, such as novel drug delivery strategies and

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nanostructured material design, hold promise in reducing incidences of infection, encrustation and blockage, and therefore reduction in patient morbidity and mortality.

3.9

References

1 gorman s.p. and tunney, m.m. (1997) Assessment of encrustation behaviour on urinary tract biomaterials. J. Biomaterial Appl. 12(2): 136–166. 2 jones, d.s., bonner, m.c., akay, m., keane, p.f., and gorman, s.p. (1997) Sequential polyurethane-poly(methylmethacrylate) interpenetrating polymer networks as ureteral biomaterials: Mechanical properties and comparative resistance to urinary encrustation. J. Mater. Sci.: Mater. Medicine 8: 713–717. 3 gorman, s.p. and jones, d.s. (2003) Complications of urinary devices. In: Medical Implications of Biofilms, ed. Wilson, M. Cambridge University Press, pp. 136–170. 4 gorman, s.p., garvin, c.p., quigley, f., and jones, d.s. (2003) Design and validation of a dynamic flow model simulating encrustation of biomaterials in the urinary tract. J. Pharm. Pharmacol. 55: 461–468. 5 lam, j.s. and gupta, m. (2004) Update on ureteral stents. Urology 64: 9–15. 6 hamill, t.m., gilmore, b.f., jones, d.s., and gorman, s.p. (2007) Strategies for the development of the urinary catheter. Expert Rev. Med. Devices 4(2): 215–225. 7 herman, j.r. (1973) Urology: A View Through the Retrospectroscope. Hagerstown, Maryland, Harper and Row. 8 zimskind, p.d., kelter, t.r., and wilkerson, s.l. (1967) Clinical use of long-term indwelling silicone rubber ureteral splints inserted cystoscopically. J. Urol. 97: 840–844. 9 pollard, s.g. and macfarlane, r. (1988) Symptoms arising from double-J stents. J. Urol. 139(1): 37–38. 10 keane, p.f., bonner, m.c., johnston, s.r., zafia, a., and gorman, s.p. (1994) Characterization of biofilms and encrustation on ureteric stents in vitro. Br. J. Urol. 73(6): 687–691. 11 stickler, d.j., evans, a., morris, n., and hughes, g. (2002) Strategies for the control of catheter encrustation. Int. J. Antimicrob. Agents 19(6), 499–506. 12 choong, s., wood, s., fry, c., and whitfield, h. (2001) Catheter associated urinary tract infection and encrustation. Int. J. Antimicrob. Agents 17(4): 305–310. 13 stickler, d.j. and zimakoff, j. (1994) Complications of urinary tract infections associated with devices used for long-term bladder management. J. Hosp. Infect. 28: 177–194. 14 tunney, m.m., keane, p.f., jones, d.s., and gorman, s.p. (1996) Comparative assessment of ureteral stent biomaterial encrustation. Biomaterials 17(15): 1541–1546. 15 mardis, h.k. and kroeger, r.m. (1988) Ureteral stents. Materials. Urol. Clin. North Am. 15(3): 471–479. 16 tunney, m.m., jones, d.s., and gorman, s.p. (1999) Biofilm and biofilm-related encrustation of urinary tract devices. Methods Enzymol. 310: 558–566. 17 tieszer, c., reid, g., and denstedt, j. (1998) Conditioning film deposition on ureteral stents after implantation. J. Urol. 160: 876–881.

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18 santin, m., motta, a., denyer, s.p., and cannas, m. (1999) Effect of the urine conditioning film on ureteral stent encrustation and characterization of its protein composition. Biomaterials 20(13): 1245–1251. 19 donlan, r.m. and costerton, j.w. (2002) Biofilms: survival mechanisms of clinically relevant microorganisms. Clin. Microbiol. Rev. 15(2): 167–193. 20 dumanski, a.j., hedelin, h., edin-liljegren, a., beauchemin, d., and mclean, r.j. (1994) Unique ability of the Proteus mirabilis capsule to enhance mineral growth in infectious urinary calculi. Infect. Immun. 62(7): 2998–3003. 21 torzewska, a., staczek, p., and rozalski, a. (2003) Crystallization of urine mineral components may depend on the chemical nature of Proteus endotoxin polysaccharides. J. Med. Microbiol. 52(6): 471–477. 22 ceri, h., olson, m.e., stremick, c., read, r.r., morck, d., and buret, a. (1999) The Calgary Biofilm Device: new technology for rapid determination of antibiotic susceptibilities of bacterial biofilms. J. Clin. Microbiol. 37: 1771–1776. 23 joshi, h.b., stainthorpe, a., macdonagh, r.p., keeley, f.x., timoney, a.g., and barry, m.j. (2003) Indwelling ureteric stents: evaluation of symptoms, quality of life and utility. J. Urol. 169(3): 1065–1069. 24 damiano, r., oliva, a., esposito, c., de sio, m., autorino, r., and d’armiento, m. (2002) Early and late complications of double pigtail ureteral stent. Urol. Int. 69(2): 136–140. 25 kunin, c.m., douthitt, s., dancing, j., anderson, j., and moeschberger, m. (1992) The association between the use of urinary catheters and morbidity and mortality among elderly patients in nursing homes. Am. J. Epidemiol. 135(3): 291–301. 26 brosnahan, j., jull, a., and tracy, c. (2004) Types of urethral catheters for management of short-term voiding problems in hospitalised adults. Cochrane Database Syst. Rev. 1: CD004013. 27 bonner, m., keane, p.f., and gorman, s.p. (1993) Characterization and antibiotic sensitivities of isolates from urethral stent biofilm. J. Pharm. Pharmacol. 45: 1445. 28 el-faqih, s.r., shamsuddin, a.b., chakrabarti, a., atassi, r., kardar, a.h., osman, m.k., and husain, i. (1991) Polyurethane internal ureteral stents in treatment of stone patients: morbidity related to indwelling times. J. Urol. 146: 1487–1491. 29. riedl, c.r., plas, e., hubner, w.a., zimmerl, h., ulrich, w., and pfluger, h. (1999) Bacterial colonisation of ureteral stents. Eur. Urol. 36(1): 53–59. 30 paick, s.h., park, h.k., oh, s.j., and kim, h.h. (2003) Characteristics of bacterial colonization and urinary tract infection after indwelling of double-J stent. Urology 62(2): 214–217. 31 warren, j.w. (2001) Catheter-associated urinary tract infections. Int. J. Antimicrob. Agents 17(4): 299–303. 32 leone, m., garnier, f., avidan, m., and martin, c. (2004) Catheter-associated urinary tract infections in intensive care units. Microbes Infect. 6(11): 1026–1032. 33 morris, n.s. and stickler, d.j. (1998) Encrustation of indwelling urethral catheters by Proteus mirabilis biofilms growing in human urine. J. Hosp. Infect. 39(3): 227–234. 34 jones, b.d. and mobley, h.l. (1987) Genetic and biochemical diversity of ureases of Proteus, Providencia and Morganella species isolated from urinary tract infection. Infect. Immun. 55(9): 2198–2203.

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35 liedl, b. (2001) Catheter-associated urinary tract infections. Curr. Opin. Urol. 11(1): 75–79. 36 belas, r. and suvanasuthi. r. (2005) The ability of Proteus mirabilis to sense surfaces and regulate virulence gene expression involves FliL, a flagellar basal body protein. J. Bacteriol. 187(19): 6789–6803. 37 rozalski, a., sidorczyk, z., and kotelko, k. (1997) Potential virulence factors of Proteus bacilli. Microbiol. Mol. Biol. Rev. 61(1): 65–89. 38 rouprêt, m., daudon, m., hupertan, v., gattegno, b., thibault, p., and traxer, o. (2005) Can urinary stent encrustation analysis predict urinary stone composition? Urology 66(2): 246–251. 39 sabbuba, n.a., stickler, d.j., mahenthiralingam, e., painter, d.j., parkin, j., feneley, r.c. (2004) Genotyping demonstrates that the strains of Proteus mirabilis from bladder stones and catheter encrustations of patients undergoing long-term bladder catheterization are identical. J. Urol. 171(5): 1925–1928. 40 gleeson, m.j. and griffith, d.p. (1993) Struvite calculi. Br. J. Urol. 71: 503–511. 41 hedelin, h., edin-liljegren, a., and grenabo, l. (1990) E. coli and urease-induced crystallisation in urine. Scand. J. Urol. Nephrol. 24(1): 57–61. 42 stickler, d.j., howe, n.s., and winters, c. (1994) Bacterial biofilm growth on ciprofloxacin treated urethral catheters. Cell Mater. 4(4): 387–398. 43 holmes, s., cheng, c., and whitfield, h.n. (1992) The development of synthetic polymers that resist encrustation on exposure to urine. Br. J. Urol. 69: 651–655. 44 choong, s.k.s., wood, s., and whitfield, h.n. (2000) A model to quantify encrustation on ureteric stents, urethral catheters and polymers intended for urological use. Br. J. Urol. 86: 414–421. 45 griffith, d.p., musher, d.m. and itin, c. (1976) Urease: the preliminary cause of infection-induced urinary stones. Invest. Urol. 13: 346–350. 46 cox, a.j., hukins, d.w.l., davies, k.e., and irlam, j.c. (1987) An automated technique for the in vitro assessment of the susceptibility of urinary catheter materials to encrustation. Eng. Med. 16(1): 37–41. 47 gleeson, m.j., glueck, j.a., feldman, l., griffith, d.p., and noon, g.p. (1989) Comparative in vitro encrustation studies of biomaterials in human urine. ASAIO Trans. 35(3): 495–498. 48. scmitz, w., nolde, a., marklein, g., and hesse, a. (1993) In virto studies of encrustations on catheters, a model of infection stone formation. Cells Mater. 3: 1–10. 49 sarangapani, s., cavedon, k., and gage, d. (1995) An improved model for bacterial encrustation studies. J. Biomed. Mater. Res. 29: 1185–1191. 50 tunney, m.m., bonner, m.c., keane, p.f., and gorman, s.p. (1996) Development of a model for assessment of biomaterial encrustation in the upper urinary tract. Biomaterials 17: 1025–1029. 51 finlayson, b. and dubois, l. (1973) Kinetics of calcium oxalate depositions in vitro. Invest Urol. 10(6): 429–433. 52 lyon, e.s. and vermeulen, c.w. (1965) Crystallization concepts and calculogenesis. Observations on artificial oxalate concretions. Invest. Urol. 3(3): 309–320. 53 stickler, d.j., morris, n.s., and winters, c. (1999) Simple physical model to study formation and physiology of biofilms on urethral catheters. Methods Enzymol. 310: 494–501.

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54 stickler, d., morris, n., moreno, m.c., and sabbuba, n. (1998) Studies on the formation of crystalline bacterial biofilms on urethral catheters. Eur. J. Clin. Microbiol. Infect. Dis. 17(9): 649–652. 55 winters, c., stickler, d.j., howe, n.s., williams, t.j., wilkinson, n., and buckley, c.j. (1995) Some observations on the structure of encrusting biofilms of Proteus mirabilis on urethral catheters. Cells Mater. 5: 245–253. 56 stickler, d.j., jones, g.l., and russell, a.d. (2003) Control of encrustation and blockage of Foley catheters. Lancet 361: 1435–1437. 57 morris, n.s., stickler, d.j., and winters, c. (1997) Which indwelling urethral catheters resist encrustation by Proteus mirabilis biofilms? Br. J. Urol. 80: 58–63. 58 stickler, d.j., young, r., jones, g., sabbuba, n., and morris, n. (2003) Why are Foley catheters so vulnerable to encrustation by crystalline bacterial biofilm. Urol. Res. 31: 306–311. 59 chakravarti, a., gangodawiula, s., long, m.j., morris, n.s., blacklock, a.r., and stickler, d.j. (2005) An electrified catheter to resist encrustation by Proteus mirabilis biofilm. J. Urol. 174: 1129–1132. 60 jones, d.s, djokic, j., and gorman, s.p. (2006) Characterization and optimization of experimental variables within a reproducible bladder encrustation model and in vitro evaluation of the efficacy of urease inhibitors for the prevention of medical device-related encrustation. J. Biomed. Mater. Res. B. Appl. Biomater. 76(1): 1–7. 61 tunney, m.m., keane, p.f., and gorman, s.p. (1997) Assessment of urinary tract biomaterial encrustation using a modified Robbins device continuous flow model. J. Biomed. Mater. Res. (Appl. Biomater.) 38: 87–93. 62 mccoy, w.f., bryers, j.d., robbins, j., and costerton, j.w. (1981) Observations of fouling biofilm formation. Can. J. Microbiol. 27(9): 910–917. 63 ceri, h., olson, m.e., stremick, c., read, r.r., morck, d., and buret, a. (1999) The Calgary Biofilm Device: new technology for rapid determination of antibiotic susceptibilities of bacterial biofilms. J. Clin. Microbiol. 37(6): 1771–1776. 64 ceri, h., olson, m., morck, d., storey, d., read, r., buret, a., and olson, b. (2001) The MBEC Assay System: multiple equivalent biofilms for antibiotic and biocide susceptibility testing. Methods Enzymol. 337: 377–385.

4 Ureteral stents: design and materials D. L A N G E and B. H. C H E W, University of British Columbia, Canada

Abstract: Ureteral stents are hollow tubes used extensively throughout urology to facilitate drainage of the kidney and ureter. They are often used when the urinary tract becomes obstructed – typically by kidney stones. Common complications associated with indwelling ureteral stents include discomfort, infection and encrustation. New biomaterials and coatings have been used in stent design in an attempt to alleviate these stent symptoms; however, the perfect stent addressing all of these issues remains to be designed. Based on the current literature, biodegradable and drug-eluting technology appear to be the future of ureteral stent design. This chapter focuses on discussing past and present biomaterials used in ureteral stent design as well as the direction that ureteral stent design is currently taking. Key words: ureteral stent, biomaterials, ureter, stent design.

4.1

Introduction

4.1.1 The uses of stents in urology Ureteral stents are hollow tubes inserted into the kidney, ureter and bladder to provide drainage of urine from the kidney to the bladder. Ureteral stents provide upper urinary tract drainage to treat renal obstruction. Throughout urology, ureteral stenting is used for patients presenting with pain from a kidney stone, a kidney infection secondary to an obstructing stone or following endoscopic surgery such as fiberoptic ureteroscopy with intracorporeal lithotripsy, extracorporeal shockwave lithotripsy, endouretorotomy (incision for ureteral stricture disease) or endopyelotomy (incision of ureteropelvic junction obstruction). In addition to these, ureteral stenting is often used following ureteral trauma caused by endoscopic or open surgery, as well as external trauma. Prophylactic ureteral stenting is used prior to extracorporeal shockwave lithotripsy for larger stones to prevent renal obstruction by the stone pieces or to drain an infected urinary system. Lastly, ureteral stenting is a procedure used post-reconstruction or -renal transplantation to serve as a splint for a ureteral anastomosis as it heals (Fig. 4.1). 85

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4.1 Ureteral stent. The ‘pigtail’ curls at either end are designed to prevent the stent from migrating into the bladder or backwards into the kidney.

4.1.2 History of biomaterials Ureteral stents and catheters are commonly used in urology for a wide variety of indications that include reconstruction, urolithiasis, transplantation and trauma. The American Heritage Dictionary of the English Language defines ‘stent’ as ‘a device that keeps a bodily orifice or organ open during skin grafting’ or as ‘a slender, thin rod or catheter that is used to support a tubular structure following reconstruction’ (American Heritage, 2000). For an excellent historical review on the development of both the word ‘stent’ and the device, the reader is referred to Bloom et al. (1999). The term was first coined by Charles T. Stent (1807–1885), an English dentist who, in trying to design a new material to make dental impressions, found that milky sap from the Palaquium gutta tree was a soft material that hardened with time (Bloom et al., 1999). The first use of biomaterials in the urinary tract dates back to 3 BC when Erasistratus, a Greek physiologist and anatomist, used a metal S-shaped tube, which he named a ‘catheter’, (Bitschay and Brodney, 1956) to treat urinary retention. It is likely that he simply coined the term catheter and that this particular use was invented in ancient Egypt where lead and papyrus catheters were used for urinary drainage (Bitschay and Brodney, 1956). The first known documented description of the use of foreign material to provide drainage of the urinary

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tract is found around 1000 BC in the early Indian surgical textbook entitled the ‘Sushruta Samhita’ (Das, 1983; Bloom et al., 1994). This textbook contains descriptions of tubes composed of gold, silver, iron and wood smeared with ghee (liquid butter) used for urethral drainage of urine, management of urethral stricture and instillation of medication (Das, 1983; Bloom et al., 1994). Early in the first millennium, new materials used in the construction of urethral catheters included bronze and lead (Celsus, 1989). In the fourth century AD, Oribasius, physician to the Emperor Julian, described a catheter composed of treated paper (Denos, 1972). In the seventh century AD, Paul of Aegina, the last great Byzantine physician, inserted a strand of wool into the catheter lumen to act as a siphon (Adams, 1846). Near the end of the first millennium, Avicenna – a Saracen philosopher, scientist, physician and Koranic scholar – was the first to describe the use of catheters composed of soft and flexible materials (Wershub, 1970; Marino et al., 1993). Although he was known to use catheters made of rigid materials such as gold, silver, tin and lead, he suggested that the best catheters were soft and flexible catheters composed of material obtained from stiffened sea animal skins stuck together with cheese glue (Wershub, 1970; Marino et al., 1993). Urethral catheters made of rigid materials such as gold, silver and brass were still widely used in the second millennium (Borgognoni, 1960). In the seventeenth century, leading Renaissance surgeon Fabricius of Acquapendente introduced a cloth catheter impregnated with wax that was molded on a silver sound; and Covillard described his preference for catheters composed of rushes, wax or lead for their pliability (Denos, 1972). Industrialization resulted in the ability to manufacture improved surgical instruments and silk woven catheters were introduced by Pickel of Wurtzburg during this period (Thomas, 1933). He made catheters by weaving a silk cylinder over probes. Natural rubber was used by Michele Troia near the turn of the nineteenth century, although natural rubber had the disadvantage of softening at core body temperature and hardening at cooler temperatures (Murphy, 1985). Fortunately, more durable flexible catheters were made available in the 1840s after Charles Goodyear developed vulcanized rubber by treating natural rubber with lead. In 1851 Goodyear was awarded a patent for moldable hard rubber, which led to more rapid and less expensive production of catheters (Castiglioni, 1947). The most popular design was a flexible rubber catheter introduced in 1860 by Auguste Nélaton, known today as the ‘red rubber catheter’ (Castiglioni, 1947). Ultimately, flexible catheters became prevalent by the end of the nineteenth century and essentially replaced metal catheters (Marino et al., 1993). With the twentieth century came the proliferation of various synthetic polymeric rubber materials used in the construction of urinary tract catheters. The modern-day doublepigtail stent was designed to prevent migration and was first introduced in

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1978 by Finney (Finney, 2002). Additionally, there was the introduction of various specialized urethral catheters, urethral stents, ureteral stents and nephrostomy tubes, and various other biomaterials that could be exposed to the urinary tract environment. Polymeric materials used include latex rubber (polyisoprene), polyethylene, polyvinylchloride, polyurethane, silicone and various different proprietary polymers. Moreover, various coatings have been introduced to improve lubricity and decrease biofilm formation and encrustation. New coatings and stent materials have been designed to alter stent surfaces so that crystal adherence, bacterial adherence and encrustation are prevented. Softer biomaterials have been incorporated to try and improve patient comfort.

4.1.3 Problems arising from stents Stents have been associated with increased morbidity causing infection (Riedl et al., 1999), encrustation (Paick et al., 2003), hematuria (Damiano et al., 2002; Joshi et al., 2003; Joshi et al., 2005) and discomfort (Joshi et al., 2001; Joshi et al., 2002). The Ureteral Stent Symptom Questionnaire (USSQ) looks at different facets of life including sexual function, which has been shown to be negatively affected by the presence of a stent (Sighinolfi et al., 2007). In fact, ureteroscopes, intracorporeal lithotriptors and ureteroscopic techniques have improved to the point that the major cause of morbidity of procedures such as ureteroscopy has become the postoperative stent left in situ. This has prompted studies to examine differences in postoperative complications and stone-free rates in stented compared with non-stented patients following ureteroscopic lithotripsy. These studies have shown that stents are not a routine requirement following uncomplicated ureteroscopy (Hosking et al., 1999; Borboroglu et al., 2001; Denstedt et al., 2001). The most severe problem arising from stents is the ‘forgotten stent’ that is left in place for several months to several years. These stents become encrusted and create difficulty for both patient and urologist, especially since their removal involves complicated surgical procedures (Rana and Sabooh, 2007) and may result in loss of the renal unit and potentially even death (Singh et al., 2005).

4.2

Current stent biomaterials

The synthetic polymer, polyethylene, was previously used in stent construction, but was abandoned due to its stiffness, brittleness and tendency to fragment. Blends of polyethylene and other polymers, such as polyurethane, have been shown to resist encrustation (Gorman et al., 1998; Gomha et al.,

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2004). Silicone is currently the most biocompatible stent material as it is the most resistant to biofilm formation, infection and encrustation (Watterson et al., 2003b), and is one of the most lubricious materials available (Jones et al., 2004). However, its flexibility and elasticity make handling difficult, particularly over a guidewire through tortuous or tight ureters. In addition, the low tensile strength of silicone makes it susceptible to extrinsic compression. The development of stent materials aims to meld the flexibility and elasticity of silicone with the rigidity of polyethylene. This goal resulted in the development of polyurethane, the most common class of polymer currently used in stents. However, polyurethane is a stiff material that causes patient discomfort and significant ureteral ulceration and erosion have been reported in an animal model (Marx et al., 1988). New proprietary materials and combinations are softer, more comfortable and easier to maneuver within the urinary tract. Examples of commonly used materials in stents include Percuflex® (Boston Scientific Corporation, Natick, MA, USA), Silitek® (Surgitek, Medical Engineering Corporation, Racine, WI, USA) C-Flex® (Consolidated Polymer Technologies, Clearwater, FL, USA), Tecoflex® (Thermedics, Wilmington, MA, USA) and ethylene-vinyl-acetate (from the polyefin family of which polyethylene is a member). They have been designed to provide rigidity to facilitate surgeon handling and to provide adequate drainage while being soft enough to limit patient discomfort.

4.2.1 New materials New materials include metal stents that are designed to keep the ureter open in the face of extrinsic ureteral compression secondary to malignant lymphadenopathy. Ureteric obstruction may result in renal impairment, pain and infection requiring urinary diversion (Chitale et al., 2002). As these stents must remain in place for long periods of time, they need to be exchanged frequently and are more susceptible to infection and encrustation leading to obstruction. The goal in this patient population is to develop a stent that maintains ureteral patency despite extrinsic compression, yet is soft enough to minimize discomfort. Furthermore, the stent should also be resistant to encrustation and infection. Metal ureteral stents were first employed by Pauer in 1992 (Pauer and Lugmayr, 1992) and have been used in the treatment of malignant ureteric obstruction (Kulkarni and Bellamy, 2001), ureteral strictures (Daskalopoulos et al., 2001) and ureteropelvic junction obstruction (Barbalias et al., 2002a). Current problems associated with metallic stents include biofilm formation, infection, migration and tissue hyperplasia leading to luminal obstruction (Barbalias et al., 2002b).

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4.2.2 Metal ureteral stents Owing to the recurring problem of encrustation and infection leading to premature stent removal, the need for a stent that is resistant to the deposition of urinary materials has become apparent. Metal stents were originally used for the relief of end-stage malignant disease, where the ureteral stricture is caused directly by the tumor or indirectly via pressure of a tumor on the ureter. Pauer and Lugmayr (1996) used metallic ureteral stents to treat 54 malignant stenoses of the ureter in 40 patients via the implantation of a self-expandable permanent endoluminal stent (SPES), the Wallstent. During the follow-up period of 10.5 months, 51 ureters were kept sufficiently open. Of these, 51% needed no further intervention, while 49% needed intervention to re-establish patency. This is in comparison with the control patients where insufficiency of the inserted double-J catheters occurred at a mean 4.3 weeks. One of the drawbacks of these metal stents, however, is that they induce local urothelial hyperplasia, with ingrowth of tissue through the struts causing recurrent obstruction over longterm placement. Recently, a nickel–cobalt–chromium–molybdenum alloy, double-pigtailed stent (ResonanceTM stent, Cook Urological, Ireland) has been developed to provide long-term urinary drainage in patients with malignant ureteric strictures (Fig. 4.2).

4.2 ResonanceTM metal stent. This tightly coiled metal stent provides drainage around the coils and is particularly useful in patients with malignant ureteric obstruction.

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The tight winding of this stent helps to maintain stent flexibility and movement, while resisting ingrowth of tissue. In addition to this, the movement of the stent causes opening of the coils, allowing the fluid to access the lumen. Its successful use has been described in a patient with malignant ureteric obstruction secondary to metastatic breast cancer where management using two 6F polyurethane stents in the same ureter had failed (i.e. inadequate drainage was achieved) (Borin et al., 2006). In a study involving 15 patients, this stent provided adequate long-term (up to 12 months) urinary drainage in patients with malignant ureteric obstruction without significant bulky pelvic disease. These stents were also found to have minimal encrustation (Wah et al., 2007). A further study has shown that this metal stent provides its best drainage when the ureter is tightly compressed onto its outer surface. This is believed to be the result of increased flow between the coils of the metal stent due to the compression and simulates the clinical situation of ureteral compression due to malignancy. It was this feature that makes the ResonanceTM stent useful in patients in which the ureter is obstructed due to malignancy (Blaschko et al., 2007). Metal stents have also been used in the treatment of benign upper urinary tract occlusions (Li et al., 2007). Li et al. recently completed a long-term study in which 13 patients between October 1995 and December 1998 with benign upper tract occlusions were treated by the implantation of a metallic stent made of titanium–nickel alloy (Andaxing Company, Beijing, China). Over the mean follow-up time of 92 months, ureteral patency was achieved in 6 patients, while assisted patency via a double-J stent was achieved in a further 3 patients. Colonization of the stents by Pseudomonas aerugionsa and Candida albicans in two further patients led to the removal of the stents, although these patients also had a double-J stent inserted at the same time, which is believed to have been the cause of the infections. Overall these studies indicate that the use of metal stents to resolve ureteral obstruction due to metastatic cancer or benign obstructions has shown some promising results. Despite this, however, more work will need to be performed to address issues such as patient comfort.

4.2.3 Bare metal stents Bare metal stents act as self-expanding scaffolding and are used extensively in cardiology to maintain coronary artery patency. Following insertion, the lumen of the stent becomes epithelialized and integrated into the endothelium over time. Bare metal stents have also been used in the ureter to treat ureteral strictures (Liatsikos et al., 2007a). The overall 4 year patency rate was 57% in 24 ureteroileal conduits using permanent metal stents. Restenosis was a problem reported by the authors and the patients required

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either repeat dilations or repeated double-J stent changes. Re-stenosis is also a problem shared by the metal stents used in coronary artery angioplasty. Drug-eluting technology has been applied to cardiology metal stents and has become the new focus for stents in urology. Drug-eluting stents are discussed in Section 4.5 below.

4.2.4 Biodegradable stents Despite the fact that biocompatible materials and stent designs have improved over the years, they have one key disadvantage: the fact that they have to be removed endourologically. More recent research has focused on the design of stents that do not need to be removed and are biodegradable. The design of a biodegradable stent must take into consideration the biocompatibility properties of the material, as well as the degree of expansion and degradation rates, and most importantly it must be able to fulfill the basic requirement of a stent in that it must be able to guarantee urinary flow from the renal pelvis through the ureter and into the bladder for the desired period of time. One of the challenges involved in designing a biodegradable stent is the control of the rate of degradation. Schlick and Planz (1997, 1998) designed a stent composed of plastics, the degradation of which was dependent upon the urine pH. In vitro experiments with artificial urine showed that the stents were stable at urine pH of less than 7.0 for at least 30 days, while they dissolved completely within 48 hours at a pH greater or equal to 7.0. The principle behind this stent is that it will remain stable at the physiological urine pH of 5–6, and can be triggered to dissolve by medically altering urinary pH. Although very promising, this technology remains at an experimental stage and awaits animal trials. An additional factor that may need to be taken into consideration is the influence of encrustation and protein deposition, as it can form a platform for bacterial adherence and infection. In addition to this, stent encrustation and bacterial adherence may also influence urine pH. Uropathogens in general are known to increase urine pH, and may have an effect in vivo. In addition to this, medically increasing the urinary pH may introduce an additional risk for infection as a more alkaline pH favors bacterial survival and increased stone formation (calcium phosphate and struvite stones). In addition, encrustation of the stent may prevent its exposure to the environment and alkaline pH, thus limiting its rate of decomposition. A spiral stent (Spirastent®, Urosurge Medical, Coralville, IA, USA) is a polyurethane stent with metal helical ridges designed to prevent kinking and compression in chronically obstructed patients. In in vitro studies, this stent increased flow and theoretically increased the space between stent and ureter to facilitate passage of stone fragments (Stoller et al., 2000). The

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spiral design has been incorporated into biodegradable materials for urethral stents (Isotalo et al., 2002; Laaksovirta et al., 2002). Laaksovirta et al. (2002) used a self-reinforced poly-l-lactic and poly-l-glycolic acid (SRPLGA) copolymer spiral urethral stent (SpiroFlow® stent, Bionx Implants Ltd, Tampere, Finland) following prostatic laser coagulation. This stent degraded in 6–8 weeks and resisted encrustation at 4 weeks in artificial urine. After 8 weeks, the SpiroFlow® stent was significantly less encrusted than the metal urethral stents Prostakath® (Engineers and Doctors A/S, Copenhagen, Denmark) and Memokath® 028 (Engineers and Doctors A/S). SR-PLGA is the most commonly utilized material for prostatic stents, but it has also been developed as a ureteral stent and may be incorporated into new degradable, encrustation-resistant ureteral stents in the future (Olweny et al., 2002). A biodegradable stent designed to provide drainage for 48 hours was previously available. Lingeman et al. (2003) reported a Phase I trial on the temporary ureteral drainage stents (TUDS®, Boston Scientific) in 18 patients who underwent percutaneous nephrolithotomy. Nephrostograms at 48 hours after insertion showed good flow and all stents were dissolved after 1 month. Further studies showed less promising results, however; a Phase II trial showed satisfactory drainage and degradation in only 68 of 87 patients implanted with a TUDS® after ureteroscopy. Early stent extrusion occurred in 17 patients (less than 48 hours). In 3 patients, stent fragments were retained for more than 3 months, requiring shockwave lithotripsy and ureteroscopy to clear. Currently, this stent is commercially unavailable. A new biodegradable stent is also being developed. We are currently involved in developing a new biodegradable stent (Indevus Pharmaceuticals, Lexington, MA, USA) that dissolves within 2–4 weeks using an animal model. Clinical trials are scheduled to begin in late 2008. Properties such as stent softness of these biodegradable stents may improve patient comfort. Stents typically require a secondary cystoscopy for their removal, unless a suture tether is used to remove the stent. Avoiding a secondary procedure for ureteral stent removal would decrease patient morbidity and make this technology attractive.

4.3

Stent coatings

One of the most common stent coatings is hydrogel, which consists of hydrophilic polymers that absorb water and increase lubricity and elasticity (Marmieri et al., 1996). These properties facilitate stent placement, making the device rigid and easily maneuverable in its dry state. Once exposed to urine, hydrogel begins to absorb and trap water in its polyanionic structure, causing it to soften and potentially increase patient comfort. Data on

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encrustation and infection are less convincing, as hydrogel has been shown to both reduce (Gorman et al., 1998) and increase (Desgrandchamps et al., 1997) encrustation and biofilm formation. Glycosaminoglycan (GAG), a normal constituent of urine, is a natural inhibitor of crystallization. Other novel stent coatings include: pentosan polysulfate (Zupkas et al., 2000) (a member of the glycosaminoglycan family, a normal constituent of urine and a natural inhibitor of crystallization); phosphorylcholine (Stickler et al., 2002) (a constituent of human erythrocytes that mimics a natural lipid membrane); and polyvinyl pyrrolidone (Tunney and Gorman, 2002) (a hydrophilic coating, similar to hydrogel, that absorbs water). Attempts to reduce encrustation have involved the use of other stent coatings, such as the bacterial enzyme oxalate decarboxylase, which has been shown to decrease encrustation in silicone disks placed in rabbit bladders (Watterson et al., 2003a).A novel coating of monomethoxy poly(ethylene glycol)-3, 4-dihydroxyphenylalanine (mPEG-DOPA3), a natural constituent produced by mussels that produces strong adhesive properties, also has the ability to avoid biofouling in the environment. The polyethyelene (PEG) component provides the antifouling property while the DOPA3 provides the adherence that PEG lacks on its own. Adherence of these combined compounds on silicone disks has resulted in a strong ability to resist bacterial adherence and growth in vitro (Ko, 2007). More recently, plasma-deposited diamond-like carbon coatings have been used to coat stents in an attempt to prevent encrustation (Laube et al., 2007). In vitro experiments have shown a 30% decrease in encrustation of these stents in artificial urine compared with the non-coated controls. Ongoing clinical trials appear to indicate a further enhancement of these results in vivo; however, a mechanism for this needs to be elucidated. Encrustation of ureteral stents remains one of the most common problems associated with ureteral stenting and more research will need to be done in order to achieve an optimal stent design which resists the deposition of bacteria, minerals and proteins. In vascular medicine, the coagulant heparin has been shown to inhibit bacterial attachment to venuous catheters (Ruggieri et al., 1987; Appelgren et al., 1996), and this has been attributed to its highly negative charge. Similarly, effects of heparin have also been observed for ureteral stents. Riedl et al. (2002) used heparin-coated and uncoated polyurethane ureteral stents and inserted them into obstructed ureters with indwelling times between 2 and 6 weeks. Electron microscopy showed that the uncoated control stents were covered with amorphous inorganic deposits and bacterioal biofilms as early as 2 weeks following stent insertion, while the heparin-coated stents remained unaffected by encrustation following 6 weeks of indwelling time. More recently, Cauda et al. (2008) performed

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a long-term study involving patients with bilateral ureteral obstructions treated via the insertion of a heparin-coated stent into one ureter, and an uncoated control stent into the other ureter. Overall, the uncoated control stents were found to be encrusted with amorphous, crystalline inorganic deposits and bacterial biofilm as early as 1 month post-insertion, while the heparin-coated stents remained visibly free of encrustation as long as 10 months post-insertion. Biofilm encrustation was evident only on the external surface of the coated stent after 1 year of being in place. These in vivo results are quite promising, although clinical trials involving a larger number of patients are needed to ensure that the heparin coating is effective across a broader patient range.

4.4

Stent design

Stent design is the most important factor determining whether a stent is good at providing efficient drainage of urine without further complications such as encrustation and patient discomfort. Much work has gone into designing the ‘perfect’ stent. The path taken by urine in a stented ureter is quite complex, as it passes through both the center of the stent and also the space between the stent wall and the ureter itself. Over the years the design has changed to include holes in the walls of stents to facilitate urine movement through the stent. In addition to this, the pigtail ends of stents, which hold them in place, provide further complexity to the urine flow, as these tails contain several ports for fluid in- and outflow. Recent work performed by Tong et al. (2007) has quantitatively analyzed the urine flow through a stented ureter. The results revealed that in the absence of blockages, the urine path did not include any of the passthrough holes found in the stent wall. In contrast, when a blockage was present due to encrustation or biofouling, urine was found to flow through the holes and adjacent to the blockage. Furthermore, only the tail port that is closest to the stent proper was involved in urine flow. Altogether, this simulation method gave insights into the way that stent design affects functionality and will be a great tool in determining the effectiveness of newer stent designs. Although the exact cause of stent-related symptoms is not fully understood, one hypothesis is that the bladder urothelium becomes irritated by the distal stent curl. In an attempt to improve comfort, a dual-durometer stent was designed using two different materials – the renal curl is composed of a more rigid material to maintain the renal coil and prevent distal migration; the bladder curl is much softer, maximizing patient comfort. Examples include the Sof-Curl® (ACMI Corporation, Southborough, MA, USA), Polaris® (Boston Scientific) and Loop Polaris® (Boston Scientific). The tail stent (Percuflex Tail Plus® stent, Boston Scientific) has a proximal

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4.3 Open-Pass stent designed to remove stone fragments inside each basket.

7F lumen which tapers to a soft, lumenless 3F tail that does not coil in the bladder (Dunn et al., 2000). The tail stent has been reported to be less irritating to the bladder than conventional double-pigtail stents (Liatsikos et al., 2002). A Loop Polaris® stent has a regular pigtail as the renal retention portion and two soft loops in the bladder to prevent migration, decrease discomfort and facilitate cystoscopic removal. To date, no clinical studies have been performed using this stent. Other stents have been designed in an attempt to facilitate passage of fragments following lithotripsy. The Open-Pass stent (Fossa Medical, Milton, MA, USA) is a self-expanding ureteral stent which was designed to cause dilatation of the ureter for improved urinary flow and stone passage (L’Esperance et al., 2007) (Fig. 4.3). This stent also contains expanding baskets along the sides, which serve to trap any small stone fragments to facilitate extraction when the stent is removed. In a clinical study involving 42 patients, this stent was found to allow for passage of stones while it is in place, as well as the removal of stones via the baskets upon its extraction. All patients underwent stent removal without complication, indicating that injury due to stone entrapment by the baskets and movement of the stones around the stent are not a concern. Malignant ureteric obstruction can cause significant problems with upper tract drainage and standard material stents may not be able to overcome the terrific extrinsic force on the ureter to provide adequate drainage. Until recently the only solutions were to insert either a nephrostomy tube or two double-J stents to facilitate drainage (Rotariu et al., 2001). A novel stent design utilizes the premise of two side-by-side J stents. The Open-I and Open-8 stents (Fossa Medical) result in more internal luminal flow due to their design (Fig. 4.4). The Open-8 is designed to maintain intraluminal flow despite extrinsic compression (Fig. 4.5). Its proprietary antimicrobial coating provides up to 1 year of indwelling time.

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4.4 The ‘I-beam’ configuration of Open-8 and Open-I ureteral stents. This mimicks putting together two double-J ureteral stents and provides flow both in and through the lumen of the stent.

Open-8

Percuflex

Silicone

4.5 Flow areas for 7F stents. The ‘I-beam’ configuration of the Open-8 and Open-I ureteral stents provides excellent flow through the lumen of the stent (shaded area).

4.5

Drug-eluting stents

The most serious complications of long-term stenting involve infection triggered by bacterial adherence and biofilm formation on the surfaces of stents as well as patient discomfort due to stent placement. Much research has gone into the prevention of infection, and the most promising results have come from drug-eluting stents. Triclosan is an antimicrobial used in many products including soap, surgical scrub, toothpaste and mouthwash. It inhibits the highly conserved bacterial enoyl-acyl carrier protein reductase, which is responsible for fatty acid synthesis and cell growth. Cadieux et al. (2006) reported that, compared with control stents, triclosan-loaded

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stents implanted in rabbit bladders infected with Proteus mirabilis were associated with significantly fewer urinary tract infections. Chew et al. (2006) have shown that bacterial adherence to triclosan-eluting stents is markedly reduced compared with regular stents. These studies indicate that human clinical trials involving these stents are warranted. Ureteral stents may also be loaded with pharmaceuticals to aid patient comfort, and to prevent encrustation. Irritative and painful stent symptoms have traditionally been managed with oral medications such as anticholinergics and analgesics, or even by stent removal. Drug-eluting stents release a medication that acts locally on the bladder to decrease irritation and pain. In an attempt to determine which medication might improve stent-related symptoms, Beiko et al. (2004) intravesically instilled three different medications into the bladders of 40 patients who had received ureteral stents at the time of shockwave lithotripsy. Intravesical ketorolac significantly reduced flank pain scores following stent insertion compared with lidocaine or oxybutynin following shockwave lithotripsy. Liatsikos et al. (2007b) have tested paclitaxel-eluting metal stents in the pig ureter to examine the tissue effects and stricture formation. Paclitaxel-eluting stents produced less ureteral inflammation and hyperplasia of the surrounding tissue compared with the bare metal stents. Ureteral patency was lost in the control stents and maintained by the paclitaxel-eluting stents. These studies were carried out over a 21-day period and require further validation via long-term animal trials. Stent encrustation worsens with increased indwelling time and concurrent infection with urease-producing organisms. Oxalate is normally broken down in the gastrointestinal tract by the enzyme oxalate decarboxylase, which is found in a commensal organism Oxalobacter formigenes. Oxalate that escapes degradation and fecal excretion is absorbed into the bloodstream and filtered in the kidneys where, under certain conditions, it can combine with calcium to form calcium oxalate stones. Watterson et al. (2003a) coated silicone disks with oxalate decarboxylase and implanted these into rabbit bladders. After 30 days, the oxalate decarboxylase-coated disks were significantly less encrusted than control disks. Coating ureteral stents with such an enzyme could theoretically prevent encrustation as the stent would elute an enzyme to degrade urinary oxalate.

4.6

Conclusions and future trends

Urinary biomaterials continue to improve. Newer stent designs with softer materials and better coatings will improve patient tolerability of ureteral stents. Three key areas need to be addressed in improving the use of stents: comfort, infection and encrustation. Studies aimed at determining the effectiveness of drug-eluting stents in addressing common problems such as

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encrustation, infection and discomfort show very promising results, and stent design will likely focus on this developing technology. Drugs inserted directly into the stent compound or coated on the stent to release over time may help with common problems such as pain, infection and encrustation. Other drugs may also be used to help with kidney stone passage, relief of obstruction, or to treat upper tract urothelial cancer. Further research will also be focused on biodegradable stent materials that may improve comfort and avoid the need for a secondary cystoscopic procedure to remove the stent. Softer biomaterials have already been designed and continue to improve, although more work will need to be done, as these materials were shown not to alleviate stent symptoms (Lennon et al., 1995).

4.7

References

adams, f. (1846) The Seven Books of Paulus Aegineta, London, Sydenham Society. american heritage (2000) The American Heritage Dictionary of the English Language, Boston, Houghton Mifflin Company. appelgren, p., ransjo, u., bindslev, l., espersen, f. and larm, o. (1996) Surface heparinization of central venous catheters reduces microbial colonization in vitro and in vivo: results from a prospective, randomized trial. Crit Care Med, 24, 1482–9. barbalias, g. a., liatsikos, e. n., kagadis, g. c., karnabatidis, d., kalogeropoulou, c., nikiforidis, g. and siablis, d. (2002a) Ureteropelvic junction obstruction: an innovative approach combining metallic stenting and virtual endoscopy. J Urol, 168, 2383–6; discussion, 2386. barbalias, g. a., liatsikos, e. n., kalogeropoulou, c., karnabatidis, d., zabakis, p., athanasopoulos, a., perimenis, p. and siablis, d. (2002b) Externally coated ureteral metallic stents: an unfavorable clinical experience. Eur Urol, 42, 276–80. beiko, d. t., watterson, j. d., knudsen, b. e., nott, l., pautler, s. e., brock, g. b., razvi, h. and denstedt, j. d. (2004) Double-blind randomized controlled trial assessing the safety and efficacy of intravesical agents for ureteral stent symptoms after extracorporeal shockwave lithotripsy. J Endourol, 18, 723–30. bitschay, j. and brodney, m. l. (1956) A History of Urology in Egypt, New York, Riverside Press. blaschko, s. d., deane, l. a., krebs, a., abdelshehid, c. s., khan, f., borin, j., nguyen, a., mcdougall, e. m. and clayman, r. v. (2007) In-vivo evaluation of flow characteristics of novel metal ureteral stent. J Endourol, 21, 780–3. bloom, d. a., clayman, r. v. and mcdougal, e. (1999) Stents and related terms: a brief history. Urology, 54, 767–71. bloom, d. a., mcguire, e. j. and lapides, j. (1994) A brief history of urethral catheterization. J Urol., 151, 317–25. borboroglu, p. g., amling, c. l., schenkman, n. s., monga, m., ward, j. f., piper, n. y., bishoff, j. t. and kane, c. j. (2001) Ureteral stenting after ureteroscopy for distal ureteral calculi: a multi-institutional prospective randomized controlled study assessing pain, outcomes and complications. J Urol, 166, 1651–7. borgognoni, t. (1960) The Surgery of Theodoric, New York, Appleton-CenturyCrofts, Inc.

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borin, j. f., melamud, o. and clayman, r. v. (2006) Initial experience with full-length metal stent to relieve malignant ureteral obstruction. J Endourol, 20, 300–4. cadieux, p. a., chew, b. h., knudsen, b. e., dejong, k., rowe, e., reid, g. and denstedt, j. d. (2006) Triclosan loaded ureteral stents decrease proteus mirabilis 296 infection in a rabbit urinary tract infection model. J Urol, 175, 2331–5. castiglioni, a. (1947) A History of Medicine, New York, Knopf. cauda, f., cauda, v., fiori, c., onida, b. and garrone, e. (2008) Heparin coating on ureteral Double J stents prevents encrustations: an in vivo case study. J Endourol, 22, 465–72. celsus, a. c. (1989) Classics of Medicine Library, Birmingham, De Medicina. chew, b. h., cadieux, p. a., reid, g. and denstedt, j. d. (2006) In-vitro activity of triclosan-eluting ureteral stents against common bacterial uropathogens. J Endourol, 20, 949–58. chitale, s. v., scott-barrett, s., ho, e. t. and burgess, n. a. (2002) The management of ureteric obstruction secondary to malignant pelvic disease. Clin Radiol, 57, 1118–21. damiano, r., oliva, a., esposito, c., de sio, m., autorino, r. and d’armiento, m. (2002) Early and late complications of double pigtail ureteral stent. Urol Int, 69, 136–40. das, s. (1983) Shusruta of India, the pioneer in the treatment of urethral stricture. Surg Gynecol Obstet, 157, 581–2. daskalopoulos, g., hatzidakis, a., triantafyllou, t., delakas, d., anezinis, p., metaxari, m. and cranidis, a. (2001) Intraureteral metallic endoprosthesis in the treatment of ureteral strictures. Eur J Radiol, 39, 194–200. denos, e. (1972) From the Renaissance to the nineteenth century. In Murphy, L. J. T. (Ed.), The History of Urology, Springfield, Charles C. Thomas. denstedt, j. d., wollin, t. a., sofer, m., nott, l., weir, m. and d’a honey, r. j. (2001) A prospective randomized controlled trial comparing nonstented versus stented ureteroscopic lithotripsy. J Urol, 165, 1419–22. desgrandchamps, f., moulinier, f., daudon, m., teillac, p. and le duc, a. (1997) An in vitro comparison of urease-induced encrustation of JJ stents in human urine. Br J Urol, 79, 24–7. dunn, m. d., portis, a. j., kahn, s. a., yan, y., shalhav, a. l., elbahnasy, a. m., bercowsky, e., hoenig, d. m., wolf, j. s., jr, mcdougall, e. m. and clayman, r. v. (2000) Clinical effectiveness of new stent design: randomized single-blind comparison of tail and double-pigtail stents. J Endourol, 14, 195–202. finney, r. p. (2002) Experience with new double J ureteral catheter stent. 1978. J Urol, 167, 1135–8; discussion, 1139. gomha, m. a., sheir, k. z., showky, s., abdel-khalek, m., mokhtar, a. a. and madbouly, k. (2004) Can we improve the prediction of stone-free status after extracorporeal shock wave lithotripsy for ureteral stones? A neural network or a statistical model? J Urol, 172, 175–9. gorman, s. p., tunney, m. m., keane, p. f., van bladel, k. and bley, b. (1998) Characterization and assessment of a novel poly(ethylene oxide)/polyurethane composite hydrogel (Aquavene) as a ureteral stent biomaterial. J Biomed Mater Res, 39, 642–9. hosking, d. h., mccolm, s. e. and smith, w. e. (1999) Is stenting following ureteroscopy for removal of distal ureteral calculi necessary? J Urol, 161, 48–50.

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isotalo, t., talja, m., valimaa, t., tormala, p. and tammela, t. l. (2002) A bioabsorbable self-expandable, self-reinforced poly-L-lactic acid urethral stent for recurrent urethral strictures: long-term results. J Endourol, 16, 759–62. jones, d. s., garvin, c. p. and gorman, s. p. (2004) Relationship between biomedical catheter surface properties and lubricity as determined using textural analysis and multiple regression analysis. Biomaterials, 25, 1421–8. joshi, h. b., chitale, s. v., nagarajan, m., irving, s. o., browning, a. j., biyani, c. s. and burgess, n. a. (2005) A prospective randomized single-blind comparison of ureteral stents composed of firm and soft polymer. J Urol, 174, 2303–6. joshi, h. b., newns, n., stainthorpe, a., macdonagh, r. p., keeley, f. x., jr and timoney, a. g. (2003) Ureteral stent symptom questionnaire: development and validation of a multidimensional quality of life measure. J Urol, 169, 1060–4. joshi, h. b., okeke, a., newns, n., keeley, f. x., jr and timoney, a. g. (2002) Characterization of urinary symptoms in patients with ureteral stents. Urology, 59, 511–16. joshi, h. b., stainthorpe, a., keeley, f. x., jr, macdonagh, r. and timoney, a. g. (2001) Indwelling ureteral stents: evaluation of quality of life to aid outcome analysis. J Endourol, 15, 151–4. ko, r., cadieux, p. a., dalsin, j. l., lee, b. p., elwood, c. n. and razvi, h. (2007) Novel uropathogen-resistant coatings inspired by marine mussels. J. Endourol., 21, A5. kulkarni, r. and bellamy, e. (2001) Nickel-titanium shape memory alloy Memokath 051 ureteral stent for managing long-term ureteral obstruction: 4-year experience. J Urol, 166, 1750–4. laaksovirta, s., isotalo, t., talja, m., valimaa, t., tormala, p. and tammela, t. l. (2002) Interstitial laser coagulation and biodegradable self-expandable, self-reinforced poly-L-lactic and poly-L-glycolic copolymer spiral stent in the treatment of benign prostatic enlargement. J Endourol, 16, 311–15. laube, n., kleinen, l., bradenahl, j. and meissner, a. (2007) Diamond-like carbon coatings on ureteral stents – a new strategy for decreasing the formation of crystalline bacterial biofilms? J Urol, 177, 1923–7. lennon, g. m., thornhill, j. a., sweeney, p. a., grainger, r., mcdermott, t. e. and butler, m. r. (1995) ‘Firm’ versus ‘soft’ double pigtail ureteric stents: a randomised blind comparative trial. Eur Urol, 28, 1–5. l’esperance, j. o., rickner, t., hoenig, d., bamberger, m. and albala, d. m. (2007) Ureteral expanding stent: a new device for urolithiasis. J Endourol, 21, 533–7. li, x., he, z., yuan, j., zeng, g., he, y. and lei, m. (2007) Long-term results of permanent metallic stent implantation in the treatment of benign upper urinary tract occlusion. Int J Urol, 14, 693–8. liatsikos, e. n., hom, d., dinlenc, c. z., kapoor, r., alexianu, m., yohannes, p. and smith, a. d. (2002) Tail stent versus re-entry tube: a randomized comparison after percutaneous stone extraction. Urology, 59, 15–19. liatsikos, e. n., kagadis, g. c., karnabatidis, d., katsanos, k., papathanassiou, z., constantinides, c., perimenis, p., nikiforidis, g. c., stolzenburg, j. u. and siablis, d. (2007a) Application of self-expandable metal stents for ureteroileal anastomotic strictures: long-term results. J Urol, 178, 169–73. liatsikos, e. n., karnabatidis, d., kagadis, g. c., rokkas, k., constantinides, c., christeas, n., flaris, n., voudoukis, t., scopa, c. d., perimenis, p., filos, k. s.,

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nikiforidis, g. c., stolzenburg, j. u. and siablis, d. (2007b) Application of paclitaxel-eluting metal mesh stents within the pig ureter: an experimental study. Eur Urol, 51, 217–23. lingeman, j. e., preminger, g. m., berger, y., denstedt, j. d., goldstone, l., segura, j. w., auge, b. k., watterson, j. d. and kuo, r. l. (2003) Use of a temporary ureteral drainage stent after uncomplicated ureteroscopy: results from a phase II clinical trial. J Urol, 169, 1682–8. marino, r. a., mooppan, u. m. and kim, h. (1993) History of urethral catheters and their balloons: drainage, anchorage, dilation and hemostasis. J Endourol, 7, 89–92. marmieri, g., pettenati, m., cassinelli, c. and morra, m. (1996) Evaluation of slipperiness of catheter surfaces. J Biomed Mater Res, 33, 29–33. marx, m., bettmann, m. a., bridge, s., brodsky, g., boxt, l. m. and richie, j. p. (1988) The effects of various indwelling ureteral catheter materials on the normal canine ureter. J Urol, 139, 180–5. murphy, l. j. t. (1985) History of Urology, Springfield, IL, Charles C. Thomas. olweny, e. o., landman, j., andreoni, c., collyer, w., kerbl, k., onciu, m., valimaa, t. and clayman, r. v. (2002) Evaluation of the use of a biodegradable ureteral stent after retrograde endopyelotomy in a porcine model. J Urol, 167, 2198–202. paick, s. h., park, h. k., oh, s. j. and kim, h. h. (2003) Characteristics of bacterial colonization and urinary tract infection after indwelling of double-J ureteral stent. Urology, 62, 214–17. pauer, w. and lugmayr, h. (1992) Metallic wallstents: a new therapy for extrinsic ureteral obstruction. J Urol, 148, 281–4. pauer, w. and lugmayr, h. (1996) [Self-expanding permanent endoluminal stents in the ureter. 5 years results and critical evaluation]. Urologe A, 35, 485–9. rana, a. m. and sabooh, a. (2007) Management strategies and results for severely encrusted retained ureteral stents. J Endourol, 21, 628–32. riedl, c. r., plas, e., hubner, w. a., zimmerl, h., ulrich, w. and pfluger, h. (1999) Bacterial colonization of ureteral stents. Eur Urol, 36, 53–9. riedl, c. r., witkowski, m., plas, e. and pflueger, h. (2002) Heparin coating reduces encrustation of ureteral stents: a preliminary report. Int J Antimicrob Agents, 19, 507–10. rotariu, p., yohannes, p., alexianu, m., rosner, d., lee, b. r., lucan, m. and smith, a. d. (2001) Management of malignant extrinsic compression of the ureter by simultaneous placement of two ipsilateral ureteral stents. J Endourol, 15, 979–83. ruggieri, m. r., hanno, p. m. and levin, r. m. (1987) Reduction of bacterial adherence to catheter surface with heparin. J Urol, 138, 423–6. schlick, r. w. and planz, k. (1997) Potentially useful materials for biodegradable ureteric stents. Br J Urol, 80, 908–10. schlick, r. w. and planz, k. (1998) In vitro results with special plastics for biodegradable endoureteral stents. J Endourol, 12, 451–5. sighinolfi, m. c., micali, s., de stefani, s., mofferdin, a., grande, a., giacometti, m., ferrari, n., rivalta, m. and bianchi, g. (2007) Indwelling ureteral stents and sexual health: a prospective, multivariate analysis. J Urol, 178, 229–31. singh, v., srinivastava, a., kapoor, r. and kumar, a. (2005) Can the complicated forgotten indwelling ureteric stents be lethal? Int Urol Nephrol, 37, 541–6. stickler, d. j., evans, a., morris, n. and hughes, g. (2002) Strategies for the control of catheter encrustation. Int J Antimicrob Agents, 19, 499–506.

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stoller, m. l., schwartz, b. f., frigstad, j. r., norris, l., park, j. b. and magliochetti, m. j. (2000) An in vitro assessment of the flow characteristics of spiral-ridged and smooth-walled JJ ureteric stents. BJU Int, 85, 628–31. thomas, g. j. (1933) Urological instruments. In Ballenger, E. G., Frontz, W. A., Hamer, H. G. and Lewis, B. (Eds), History of Urology, Baltimore, MD, Williams and Wilkins Co. tong, j. c., sparrow, e. m. and abraham, j. p. (2007) Numerical simulation of the urine flow in a stented ureter. J Biomech Eng, 129, 187–92. tunney, m. m. and gorman, s. p. (2002) Evaluation of a poly(vinyl pyrollidone)coated biomaterial for urological use. Biomaterials, 23, 4601–8. wah, t. m., irving, h. c. and cartledge, j. (2007) Initial experience with the resonance metallic stent for antegrade ureteric stenting. Cardiovasc Intervent Radiol, 30, 705–10. watterson, j. d., cadieux, p. a., beiko, d. t., cook, a. j., burton, j. p., harbottle, r. r., lee, c., rowe, e., sidhu, h., reid, g. and denstedt, j. d. (2003a) Oxalate-degrading enzymes from Oxalobacter formigenes: a novel device coating to reduce urinary tract biomaterial-related encrustation. J Endourol, 17, 269–74. watterson, j. d., cadieux, p. a., stickler, d., reid, g. and denstedt, j. d. (2003b) Swarming of Proteus mirabilis over ureteral stents: a comparative assessment. J Endourol, 17, 523–7. wershub, l. p. (1970) Urologic surgery from antiquity to the 20th century, part III. In Urology: From Antiquity to the 20th Century, St. Louis, Warren H. Greene. zupkas, p., parsons, c. l., percival, c. and monga, m. (2000) Pentosanpolysulfate coating of silicone reduces encrustation. J Endourol, 14, 483–8.

5 Metal stents in the upper urinary tract E. L I AT S I K O S, D. K A R N A BAT I D I S, P. K A L L I D O N I S and D. S I A B L I S, University of Patras, Greece

Abstract: Metal stents have been used in urology after their initial successful application in other organ systems. The indications for metal stents include malignant and benign ureteral obstruction. Several types of metal stents have been introduced and used in clinical practice with satisfactory results. Immediate as well as long-term ureteral patency can be successfully established using metal stents. Complications include encrustation, migration and endothelial hyperplasia. Future investigations will focus on the development of metal stents with fewer complications. Key words: metal stent, ureteral obstruction, ureteroileal, hyperplasia, encrustation.

5.1

Introduction

The application of metal stents (MSs) has proven successful in the vascular and biliary systems, leading several investigators to propose the expansion of their use into urology. Milroy et al. were the first to implant MSs in the urinary tract for the treatment of urethral stricture.1 In the following years, their use has been expanded to the management of benign prostate hyperplasia, urethral stricture and detrusor sphincter dyssynergia.2,3 The initial indication for implantation of MSs in the urinary tract has been in the palliative treatment of patients with end-stage malignant disease.4–9 The management of ureterointestinal strictures after urinary diversion is another application that has proven to be effective, yet challenging. Despite the promising results of the use of MSs in urology, several complications such as tissue ingrowth and recurrent obstruction could not be prevented.10–12 Several important adverse effects influence the short-term and, in particular, the long-term effectiveness of MSs inserted in stenosed ureters. Urothelial hyperplasia through the metal mesh of the stent, encrustation, infection, stent migration and interaction between stent and ureteral peristalsis are commonly encountered complications. The urological interest in MSs has focused on the development of methods and materials that could 104

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minimize and overcome the above side-effects; the ideal MS – combining radiopacity, long-term patency, low cost, and resistance to encrustation, infection and migration – is still awaited. The application of MSs in the upper urinary tract includes the implantation of MSs in the ureter in order to ensure the passage of urine from the renal pelvis to the bladder.

5.2

Types of metal stents in the upper urinary tract

The characteristics of an ideal stent probably represent an impossible goal for urological research. Nevertheless, efforts are being made to develop a stent that combines several beneficial properties. The ideal stent should be easy for the user to maneuver and have the following additional features: stability after placement, radiopacity, resistance to encrustation and infection, the ability to relieve intraluminal and extraluminal obstructions, longterm patency and low cost.5,13 The basic characteristics of the ideal ureteral stent are presented in Table 5.1.14 Five general types of MSs have been used in the ureter: self-expandable stents, balloon-expandable stents, covered stents, thermo-expandable shapememory stents and drug-eluting stents. The latter type have been used only in experimental studies in urology. Recently a new design of all-metal double-pigtail ureteral stent has been under evaluation.15 The most commonly used MSs for ureteral insertion are the self-expanding stents; nevertheless, urological research has been focused on the application of more advanced MSs, such as covered MSs, in an effort to minimize complications related to MS implantation, especially tissue ingrowth which jeop-

Table 5.1 Properties of an ideal ureteral stent Property

Description

Memory Durometer rating Elasticity Tensile strength Elongation capacity Biodurability

Maintenance of its position within the ureter Strength of memory Manipulation of its shape Crystallization and cross-linking in the biomaterials Elongation at stent breakage Ability to exist within the body without degradation of its structure and function No significant effect of the stent on the urothelium Facility of its passage or exchange Facility of stent visualization during fluoroscopy

Biocompatibility Coefficient of friction Radiopacity

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ardizes the ureteral patency. The reduction of tissue ingrowth has proven to be promising but the efficacy of covered stents was significantly limited by the high rates of migration.16 Several studies have been published by various groups, indicating the increasing interest in the field of ureteral MSs. A summary of the existing literature is presented in Table 5.2.

Table 5.2 Experience gained with various types of metal stents

Type of metal stent and investigator group Self-expanding Lugmayr and Pauer,8 selfexpanding MS Pauer and Lugmayr,9 selfexpanding MS Pollak et al.17

Lugmaur and Pauer,5 selfexpanding MS Barbalias et al.,4 selfexpanding MS and balloonexpandable MS

Number of patients/ number of ureters stented

Mean follow-up period (range)

and balloon-expandable MSs 23/30 31 weeks (3–75 weeks)

Obstruction type

M

12/15

N.A. (3–31 weeks)

M

8/11

N.A.

B, M

40/54

10.5 months (1–44 months)

M

12/14

9 months (8–16 months)

M

Description

Primary patency was 83% after 30 weeks; macrohematuria in 1 patient; encrustation in 2 patients Hematuria in 1 patient; encrustation in 2 patients and obstruction distal to the stent in 3 patients 2 of 5 malignant strictures were occluded within 1 month; 1 of 6 benign strictures remained patent in 11 months 49% reintervention and 3 ureters finally had to be abandoned

6 patients treated with self-expanding stents, 6 with balloonexpandable stents; reintervention in 3 cases due to urothelial hyperplasia, tumor ingrowth and local recurrence of primary cancer invading the upper end of the ureter

Table 5.2 Continued

Type of metal stent and investigator group 34

Number of patients/ number of ureters stented

Mean follow-up period (range)

Obstruction type

Description

Wakui et al., selfexpanding MS and balloon expandable

9/11

N.A. (3–11 months)

Barbalias et al.,10 selfexpanding MS Rapp et al.,12 selfexpanding MS Barbalias et al.18 selfexpanding MS

14/14

15 months (9–24 months)

M

2 patients required reintervention

4/6

10 months (7–12 months)

M

No recurrence

4/6

6 months (9–24 months)

M

18/24

21 months (7–50 months)

M

No restenosis in 3 patients; recurrence of stricture in one patient 2 months after implantation; no other complications observed; one patient died 12 months after stent placement Technical success 100%; Immediate poststenting success was 70.8% (17 of 24 cases); primary patency rates at 1 and 4 years were 37.8% and 22.7%, respectively Secondary intervention in 15 ureters; secondary patency rates at 1 and 4 years were 64.8% and 56.7%, respectively; periodical exchange of external– internal double-J catheters in 6 patients

16/20

8 months (6–16 months)

M

Liatsikos et al.19 selfexpanding MSs

Covered MSs Barbalias et al.16

2 ureters required insertion of additional stents (after 4 and 5 weeks) due to persistent obstruction. 2 of 3 distal stented ureters presented vesicoureteric reflux (spontaneously resolved). 2 patients died during follow-up period.

13 of 16 patients required reintervention due to stent migration; patency was 100% (follow-up end)

Table 5.2 Continued

Type of metal stent and investigator group Trueba Arguinarena and Busto20

Number of patients/ number of ureters stented 20/29

Mean follow-up period (range)

Obstruction type

N.A. (3–24 months)

B, M

3 patients underwent a second stent implantation due to migration of the first stent

B, M

3 of 15 stent migrations; a shorter Memokath 051 was inserted for prevention of bladder irritation; patency 100% (follow-up period end) 9 in 7 patients stents removed, 4 of them requiring no more intervention, 1 nephrectomy, 1 renal failure and 1 cystectomy with nephroureterectomy 22 non-functioning stents in 5 months; 10 migrated, 12 malfunctioning; 4 stents were occluded by stones after 1–10 months

Thermo-expandable shape-memory MSs 15/22 10.6 months Kulkarni and (2–21 months) Bellamy13

Kulkarni and Bellamy21

28/37

19.3 months (3–35 months)

B, M

Klarskov et al.22

33/37

14 months (3–30 months)

B, M

Resonance all-metal double – J ureteral stent 1/1 4 months Borin et al.23

M

Wah et al.15

M

15/17

M, malignant; B, benign.

2–10 months (living patients) 1 week–8 months (deceased patients)

Description

Unobstructed flow of urine 4 months after placement Stent change after 12 months in 1 patient, in 3 patients after 6 months; minimal encrustation; four patients alive at the end of the follow-up period; seven patients died with functioning stents; three failures from the outset due to bulky pelvic malignancies

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Applications of metal stents

5.3.1 Self-expandable metal stents The Wallstent (Microvasive, Natick, MA, USA) is a self-expandable endoprothesis and is composed of braided biomedical cobalt-based alloy monofilaments. Diameters range from 7 to 10 mm and the available length is 4, 6, or 8 cm. The Wallstent has been used extensively in various reports for over a decade. The Accuflex stent (Medi-tech/Boston Scientific, Boston, MA, USA) is also a self-expanding endoprosthesis composed of a monofilament wire made of titanium-based alloy (Nitinol) woven into a tubular mesh configuration. The device as a whole consists of the stent and the delivery system. Malignant and benign strictures have been managed with the insertion of either Accuflex or Wallstent MSs. The initial experience with the use of self-expandable MSs was reported by Pauer and Lugmayr in a population of 12 patients (15 obstructed ureters) in 1992.9 Candidates for the implantation of MSs were patients with ureteral obstruction due to lymph node metastases of different malignant tumors or direct tumor compression of the ureter. The follow-up time ranged between 3 and 31 weeks. Primary patency was achieved in all cases. Complications reported were a case of hemorrhage and two cases of stent encrustation. Slight obstruction due to hyperplastic regeneration of the endothelium was observed during the first 4 weeks. Full incorporation of the endoprothesis into the ureteral wall occurred 8 weeks after implantation with concurrent coverage of the stent endoluminal surface by endothelium and regression of the hyperplastic tissue. During the same year, the immediate and long-term experience with the application of the Wallstent for the management of malignant ureteral obstruction in patients fulfilling certain selection criteria was reported by the same group.8 Life expectancy of at least 6 months, current polychemoherapy, increasing levels of serum creatinine and severe clinical signs/ symptoms related to hydronephrosis were given as the selection criteria for the management of a patient with MSs. A total number of 23 patients carrying 30 stented ureters were studied. Implantation was successful in 97% of the cases; the primary patency rate was 83% after 30 weeks of follow-up. The survival rate was 81% and 61% for the patient population after 6 and 8 months, respectively. Complications were limited to a case of macrohematuria and two cases of stent encrustation. It should be noted that a double-J catheter was always inserted in the stented ureter for a period of 4 weeks in order to prevent obstruction by endothelial hyperplasia. Flueckiger et al.7 implanted self-expandable MSs percutaneously in 10 patients (13 ureters) and VanSonnenberg et al.6 implanted MSs in six patients with malignant ureteral obstruction that could not be treated with

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double-J stent placement. Primary patency was achieved in all cases of the former group while inadequate urine drainage was observed in two patients of the latter group, requiring the co-axial insertion of a double-J catheter. Urothelial hyperplasia was reported to be a complication by Flueckiger et al.7 One ureter was occluded by hyperplastic tissue 4 weeks after implantation and a double-J stent was placed. The rest of the implanted endoprotheses did not show signs of obstruction during the available follow-up time, ranging from 3 to 14 months (average 5.8 months). In conjunction with the clinical experience with MSs in the ureter, an experimental study by Wright et al.24 proposed the use of MSs as ureteral occlusion devices. Self-expanding and balloon-expandable stents were placed in both normal and stenotic canine ureters. The stents were occluded at 4 weeks after implantation regardless of the diameter of the stents used and the presence or absence of pathology in the ureters. Only one case of migration was observed. An experimental study by Desgrandchamps et al.25 investigated the efficiency of self-expanding stents in the pig ureter. Distal narrowing of the ureter was observed owing to a functional discrepancy between the adynamic stented ureter and the normal underlying ureter. Urothelial hyperplasia was also observed without playing an important role in the obstructing process. Thijssen et al.26 inserted self-expandable MSs in histologically normal canine ureters and reported that the ureters remained patent whereas the MSs did not incorporate within the ureteral wall. In addition, the epithelium/submucosa layers were expanded between wire struts. Areas of fibrosis in the submucosal layer were also present. Poor surgical candidates with both benign and malignant ureteral strictures were managed with Wallstent endoprotheses in a study published by Pollak et al. in 1995.17 Five malignant and eight benign ureteral strictures in eight patients were included in the study. Ten lesions were strictures of ureteroenteric anastomosis and only one involved the midureter. Three stents inserted for the palliative management of malignant disease remained patent until the deaths (3–5 months after implantation) of the two patients carrying them. Two stents in a patient were occluded within a month. Only one of six stents placed for benign disease remained patent after 11 months. All occlusions in benign strictures were related to ingrowth of hyperplastic urothelium and granulation tissue. The malignant occlusions were caused by tumor ingrowth and granulation tissue. Wallstent endoprostheses were deemed ineffective in providing long-term relief in patients with benign ureteroenteric strictures.17 Mid-term results on the clinical efficacy of the implantation of selfexpandable permanent endoluminal stent for the management of malignant ureteral obstruction were studied by Lugmayr and Pauer5 in a population of 40 patients with 54 malignant stenoses. The mean follow-up time was 10.5 months (ranging between 1 and 44 months). Fifty-one stented ureters

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remained patent during the follow-up period. Reintervention was necessary in 49% of the above ureters in order to maintain patency. No major complications were observed while hydronephrosis was prevented in most of the cases. The authors concluded that the implantation of self-expandable stents is a safe and effective method for the palliative treatment of tumorassociated ureteral strictures.5 Barbalias et al.10 inserted self-expandable MSs for the treatment of gynecological cancer ureteral stricture in 14 women. Follow-up time ranged between 9 and 24 months (mean period was 15 months). Successful alleviation of ureteral obstruction was present in all cases. One case of tumor ingrowth requiring additional totally co-axial stent placement and one case of borderline tumor overgrowth were observed 6 months and 12 months after the endoprothesis placement, respectively. Urothelial reaction was observed in all cases and one case required the insertion of a double-J stent cysteoscopically for a period of 1 month. A mild ureteral narrowing was noted proximal to the stent, reported as a trumpet-like configuration that did not hinder ureteral patency and remained constant during the follow-up period. The authors concluded that the implantation of self-expandable MSs is a safe and effective method for by-passing ureteral obstruction owing to gynecological malignancies.10 An interesting combination of self-expandable MSs with co-axial doubleJ stents placed in ten ureters in ten patients with malignant ureteral obstruction was introduced by Hekimoglou et al.27 in 1996. Two and three months after the implantation, the double-J stents were removed in seven patients. Six patients developed hydronephrosis, while the remaining one patient did not demonstrate any signs of urinary obstruction. The latter (female) patient tolerated the removal of the double-J stent well until her death 5 months after stent insertion. Double-pigtail stents were again inserted in the six patients with hydronephrosis and were exchanged with new ones every 3 months. Urothelial hyperplasia and encrustation were responsible for the reduced diameter of the MS lumen. The double-J stents were never removed after the MS implantation in three patients and regular exchange every 3 months took place. Urothelial hyperplasia was observed to be present until the sixth month of the follow-up for three patients. MSs were proven to be capable of maintaining the ureteral lumen patency regardless of the hyperplastic tissue or encrustation, thus allowing the easy exchange of double-J stents. The combination of the two stent types achieved internal urinary drainage.27 Placement of self-expandable MSs for the management of ureteroileal anastomotic strictures has been investigated by several authors with interesting results.11,12,18,19,28–32 The number of patients included in the studies ranged from 1 to 18, the number of ureteroileal strictures from 1 to 24 and the mean follow-up period from 6 to 22 months. Most of the authors

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did not report recurrence of the strictures. The largest population of patients with long-term follow-up has been studied by Liatsikos et al.19 A total of 18 patients with stricture at the site of ureteroileal anastomosis were treated with MS implantation and were followed up for 21 months (range 7–50 months). The technical success of stricture crossing and stenting was 100%. The immediate post-stenting patency rate was 70.8% (17 of 24 cases) while primary patency rates at 1 and 4 years were 37.8% and 22.7%, respectively. Secondary interventions were deemed necessary in 15 ureters. Secondary patency rates were 64.8% and 56.7% at 1 and 4 years, respectively. Six ureteroileal conduits required the periodical exchange of external-internal double-J catheters until the end of the follow-up period. The MSs represented the definitive treatment for the strictures for more than half of the cases, whereas the rest of the cases allowed the uncomplicated and regular exchange of double-J catheters in retrograde fashion. As a result, the proposed less invasive treatment of ureteroileal anastomotic strictures could help to avoid surgical revision and preserve patients’ quality of life. An innovative application of self-expandable MSs was introduced by Barbalias et al.33 for the management of ureteropelvic junction (UPJ) obstruction by metal stenting. Four patients were treated with MS implantation due to recurrent UPJ obstruction after open pyeloplasty. Mean follow-up time was 16 months (range 9–24 months). Successful placement and immediate patency was achieved in all cases. One patient required additional reintervention with co-axial overlapping MSs due to ingrowth of scar tissue 2 months after stent insertion. The results were promising and the authors suggested that more extensive clinical trials should be carried out in order to validate the above application of MSs.29

5.3.2 Balloon-expandable metal stents The balloon-expandable stents used in the ureter are the Strecker stent (Boston Scientific, Natick, MA, USA) and the Palmaz–Schatz (Johnson and Johnson, Warren, PA, USA). The Strecker endoprothesis is made of tantalum monofilament wire braided in an elastic tubular mesh configuration and mounted on a 7Fr delivery catheter which is easily inserted over the guide wire. Silicone sleeves secure the stent during positioning and the release takes place after inflation of the balloon. The Palmaz–Schatz stent is a seamless tube of annealed stainless steel (316L) with staggered parallel slots etched through the wall. Initially the stent is mounted on an angioplasty balloon and the stent–balloon assembly is delivered over the guidewire through a 7Fr introducing catheter to the target site. After the deployment by inflating the balloon, the stent is expanded to its final dimensions.

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The experience with balloon-expandable MSs is limited. Barbalias et al.4 implanted balloon-expandable stents in six patients (Strecker stent) and self-expandable MSs (Accuflex stent) in another six patients (14 ureters in total) with malignant ureteral obstruction. Eleven ureters remained patent for the follow-up period which ranged between 8 and 16 months (average 9 months). Secondary intervention took place in three ureters due to obstructive urethelial hyperplasia, tumor ingrowth and local recurrence of primary cancer invading the upper end of the stent. Two patients died 2 and 10 months after the stent implantation. The aforementioned trumpet-like configuration was also observed in this series. Both types of MSs had advantages and disadvantages which should be balanced against each other when the choice for the ideal stent for each case is made. Nevertheless, both stents were proven to be safe and effective in the treatment of the malignant obstructions. Wakui et al.34 used both balloon-(Palmaz–Schatz) and self-expandable (Accuflex, Wallstent) stents for the management of malignant and benign ureteric obstruction. Eleven ureters in nine patients were stented. A balloon-expandable stent was inserted in one patient and self-expandable MSs in the remaining eight. Follow-up time ranged between 3 and 11 months. During the follow-up period, nine ureters remained patent and two required the insertion of additional stents (4 and 5 weeks after the implantation) due to persistent obstruction. Transient vesicoureteric reflux in two of three stented distal ureters resolved spontaneously within 2 months after the procedure. Ureteric patency was maintained and no complications were observed in any of the patients. Two patients died, at 3 and 5 months after implantation. MSs were proposed by the authors as a safe and effective alternative to an indwelling double-pigtail catheter or percutaneous nephrostomy.34

5.3.3 Covered metal stents A self-expanding polytetrafluoroethylene-covered nitinol stent (Hemobahn Endoprosthesis, W.L. Gore and Associates, Flagstaff, AZ, USA) was implanted by Trueba Arguinarena and Busto20 for the management of both benign and malignant ureteral obstruction. Twenty patients with a total of 29 ureteral stenoses of both malignant and benign etiology were treated with the insertion of the nitinol-covered stents. The efficacy of the stent as a safe and effective treatment for ureteral stenosis was proven. In addition, high resistance to calcification was also demonstrated. Hyperplastic tissue was observed at the stent ends in only a few cases without causing any obstruction. Migration proved to be limited occurring in 3 out of 20 patients (4 stents in total). The investigators concluded that long-term follow-up is necessary to confirm the efficacy of the stent in inhibiting ureteral hyperplasia.20

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The Passager metal stent (Boston Scientific, Natick, MA, USA) is a flexible, self-expandable, nitinol stent externally coated with ultra-thin woven polyester fabric and has the additional property of shape memory. The lowporosity fabric is attached to the stent framework by polyester ligatures. The stent is presented in a compressed form in a loading cartridge. Its diameter, when fully expanded, is 8 mm and the length is 3–10 cm. The introduction of the stent is performed through long introducer sheaths. Barbalias et al.16 used the Passager stent in 16 patients (20 stented ureters) with malignant ureteral obstruction. The mean follow-up time was 6–16 months. Initiallly, all stents were successfully positioned. In 13 cases (81.2%) the stents migrated into the bladder hindering the overall ureteral patency; the mean time for migration was 1.5 months. In the remaining 3 ureters, patency was achieved without the need for further intervention. Inappropriate anchorage and increased ureteral peristalsis are the major causes of prothesis migration towards the bladder. The high rate of migration renders the use of coated MSs in the ureter an unfavorable clinical treatment.16 An interesting experimental approach for the comparison of bare MSs (Wallstent) with externally (Passager) and internally (Corvita endoluminal graft, CEG) coated stents was reported by Liatsikos et al.35. All types of stents were implanted in six pig ureters each and the follow-up period was 3 weeks. The Passager stents migrated towards the bladder in four cases. The Wallstents generated mild inflammation and metaplasia, the CEG was responsible for more pronounced inflammation in the adjacent ureter whereas the Passager stent generated severe inflammatory reaction with necrosis of the urothelium. The bare endoprothesis revealed polypoid reaction in the internal surface of the devices, a phenomenon absent in the coated stents.35 Similar results were observed by Leveilee et al.36 when implanting bare MSs and MSs lined with a porous biocompatible polymer in canine ureters. The MSs lined with the polymeric material prevented tissue ingrowth, while bare MSs caused ureteral obstruction and hydronephrosis.

5.3.4 Thermo-expandable metal stents The Memokath 051 (Engineers & Doctors A/S, Copenhagen, Denmark) is a thermo-expandable shape-memory stent that composed of nickel and titanium alloy.13 The unique tight spiral structure of the endoprothesis prevents endothelial hyperplasia. The stent softens below 10 °C and regains its shape when reheated to 50 °C. Easy removal is facilitated by the thermal shape memory. The stent has a shaft diameter of 9Fr or 10.5Fr and its wider fluted proximal end expands to 17Fr or 20Fr, respectively. Available lengths are 30, 60, 100, 150 and 200 mm. The unexpanded form of the prothesis is

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mounted over a hollow catheter with its wide end held with plastic lugs to prevent movement during insertion. The entire stent and catheter system is inserted through a 12Fr sheath. A stepped dilator is coupled with the sheath and its inner core can be removed after dilatation. The working end of the endoprothesis assembly has a port for injecting hot water to induce stent expansion. The stent has been used by Kulkarni and Bellamy13,21 for the management of both benign and malignant ureteral obstruction. Very good long-term results were demonstrated for the treatment of malignant and recurrent benign strictures. A total number of 28 cases (18 malignant and 10 recurrent benign) were managed. Mean follow-up was 19.3 months (range 3 to 35 months). At the end of the follow-up period, 15 stents were functional in 13 patients while 8 patients carrying 13 functioning stents died.21 Complications such as infection or migration were not observed.13,21 The Memokath 051 seemed to provide significant benefit over conventional double-pigtail catheters. It is noteworthy that the thermal memory permits removal, a feature unavailable to any other MS until the introduction of the Resonance double-pigtail MS. Klarskov et al.22 used the Memokath 051 for the alleviation of malignant and benign obstructions in 33 patients carrying 37 stenosed ureters. The mean follow-up period was 14 months (range 3–30 months). Ten stents migrated distally towards the bladder, 12 stents malfunctioned resulting in a total number of 22 non-functioning MSs. Stones occluded four stents in intervals after stent insertion ranging between 1 and 10 months.22

5.3.5 Drug-eluting metal stents Drug-eluting stents (DESs) attracted urological interest after their successful use in the cardiovascular system. Coronary stenting after percutaneous transluminal coronary angioplasty (PTCA) is a common practice in interventional cardiology. Neo-intimal proliferation remains a problem for the MSs implanted after coronary angioplasty, and this is also the case for ureteral MSs. In fact, tissue ingrowth is much greater in the stented vessels than it is in vessels treated with conventional PTCA alone; this could be attributed to damage of the vessel wall after stent implantation.37,38 DESs became a reality when drugs were incorporated into the polymer matrix encapsulated over the stents and it became possible to control the delivery over a period of time of drugs inhibiting cell division.39 Clinical trials revealed that DESs reduced the rate of restenosis and neo-intimal proliferation when compared with bare stents.40,41 DESs have been tested extensively using randomized trials in order to gain approval by the US Food and Drug Administration.41–44 The problems related to the use of DESs after PTCA include the potential inhibition of healing over the stent

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struts due to the drug, resulting in exposure of stent struts, and malapposition of the stents within the wall of the artery. Subacute stent thrombosis remains a possibility associated with the above suboptimal stent deployment.45 The adoption of DESs has been enthusiastic, with the rate of drugeluting stenting reaching 90% in 2006.46 In general, DESs have reduced the rate of restenosis but the incidence of late thrombosis remains higher for DESs than for bare metal stents. Stent thrombosis has a myocardial infarction rate of 70% and a mortality rate of 31–45%. No increase in myocardial infarction or number of deaths has been observed with DESs in comparison with bare MSs in simple coronary lesions. Nevertheless, DESs are mainly used in complex coronary disease where the risk of thrombosis is higher.47 In urology, DESs have only been used in experimental studies. In 2006, paclitaxel-eluting stents (TAXUS™, Boston Scientific, Natick, MA, USA) were implanted by Liatsikos et al.48 in pig ureters and resulted in reduced inflammation and smooth muscle proliferation in comparison with bare MSs (R-stent, Orbus Medical Technologies, Hoevelaken, the Netherlands). Consequently ureteral patency was maintained. The follow-up time was 21 days (3 weeks). Five of ten ureters stented with R-stents were completely occluded while two were partially stenosed. Occlusion of the DESs was not observed (Fig. 5.1). However, two DESs migrated distally within the ureter and jeopardized ureteral patency. Severe inflammation with metaplasia of the urothelium was generated by the R-stents, while polypoid reaction adherent to the internal surface of the devices was observed (Fig. 5.2). In contrast, the paclitaxel-eluting stent generated mild inflammatory response within the ureteral lumen at the site of the stent without hindering ureteral patency.48

(a)

(b)

5.1 Section of the stented pig ureteral lumen (DES). The fine epithelial lining is evident covering the metal mesh (a). The patency of the ureteral lumen is documented ureteroscopically (b).

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5.2 Section of the pig stented ureteral lumen (bare MS). Hyperplastic tissue expanding through the metal stent struts.

Shin et al.49 also applied paclitaxel-eluting, polyurethane-covered stents in a canine urethral model. The DESs exhibited less tissue hyperplasia reaction compared with the non-drug-eluting polyurethane-covered stents. These authors also proposed a design for a retrievable stent in order to make removal easier.49

5.3.6 Metal pigtail stent A recently introduced MS that shows promise is the all-metal, doublepigtail Resonance stent (Cook Ireland, Limerick, Ireland). The endoprothesis is composed of the MP35N alloy (nickel–cobalt–chromium–molybdenum alloy) which is corrosion resistant, magnetic resonance imaging (MRI) compatible and has ultra-high tensile strength. A special insertion method is required due to the fact that no end or side holes are present on the stent. The manufacturer recommends a maximum indwelling time of 12 months. Borin et al.23 successfully managed a ureteral obstruction in a 64-year-old woman with metastatic breast cancer causing retroperitoneal fibrosis with

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the insertion of a Resonance stent. Previous attempts to insert doublepigtail stents had failed. The ureters were proved to be patent during the 4 month follow-up period. Recently, the initial experience of Wah et al.15 was reported. Fifteen patients (17 ureters) with ureteral obstruction of malignant etiology were treated with the insertion of Resonance stents. One patient received a new stent 12 months after the placement of the first stent while three more patients had their stents changed at 6 months. The drainage was adequate with minimal encrustation. Four were still alive with functioning stents at the end of the follow-up period. Seven patients died with functioning stents in place (follow-up periods ranging from 1 week to 8 months). Bulky pelvic malignancies resulting in high intravesical pressure lead to exclusion of three patients from the outset. The initial experience with the Resonance MS indicated adequate long-term urinary drainage in patients with malignant ureteral obstruction without bulky pelvic disease.15 In 2007, Blaschko et al.50 carried out an experimental study to evaluate the flow in the Resonance stent comparison with the flow in a standard ureteral stent. Six Resonance stents and six standard ureteral stents were placed in six minipigs. The Resonance stent was proven to provide less overall flow than the standard ureteral stent. Nevertheless, the Resonance stent is capable of providing satisfactory drainage under extrinsic ureteral compression conditions that would be sufficient to occlude a standard stent.

5.4

Insertion techniques

Insertion techniques are different for conventional MSs and the Resonance all-metal double-pigtail MS.

5.4.1 Conventional metal stent The insertion of a conventional MS can be performed in either an antegrade or a retrograde fashion.4,51 In general, the implantation of an MS requires expertise with transurethral and percutaneous techniques within the ureteral lumen.5–7,30,52 The lengths of the MSs used for the management of ureteral obstruction range from 3 to 12 cm and the diameter is 7 mm.4,5,10,30,34,52 Initially, a standard percutaneous nephrostomy is established and a nephropyeloureterography is performed in order to localize the exact site of the obstruction. The passage through the stricture is performed with the aid of 0.018-, 0.035- or 0.038-inch guidewires depending on the pathology.4–7,10,21,31,52 Dilatation of the stenotic segment is carried out using a high-pressure balloon catheter, 8–10 mm in diameter (Fig. 5.3). Resistant strictures require repeated dilatations. Following the dilatation, the stent is inserted over the guidewire in such a position that the upper

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5.3 Balloon dilatation of the stenotic ureteral segment.

end bypasses the stricture by at least 3–4 cm, while the lower end extends intravesically for 0.5–1 cm from the ureteral orifice. If necessary, two or more stents are placed in sequence overlapping by at least 2–3 cm in order to bridge longer obstructed ureteral segments. In cases of inadequate expansion of the MS, supplementary balloon dilatations are performed.5–7,10,21,32,52 The retrograde approach for the insertion of an MS is the same in concept as the antegrade approach. After the MS implantation is completed, plain radiographs of kidneys– ureters–bladder are obtained, followed by ultrasonography and antegrade nephrotomography, to confirm ureteral patency (Fig. 5.4). The removal of the nephrostomy tube could follow. Urinalysis, urine culture, serum creatine measurement, transabdominal ultrasonography and excretory urography are performed in scheduled order. When indicated, diethylenetriamine pentaacetic acid (DTPA) renography is performed. Ureteroscopy could be deemed necessary for the assessment and treatment of eventual encrustations. Virtual endoscopy has been proposed as a non-invasive method to evaluate the patency of the stented ureter.53,54

5.4.2 Metal double-pigtail stent – Resonance The insertion technique for the Resonance stent differs from that for the conventional ureteric stents. End holes do not exist on the stent, thus the passage of the stent over a guidewire is precluded. Moreover, the flexibility

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5.4 Excretory urography: patent stented ureteral segment.

of the stent precludes the forcible pushing of the stent to negotiate a tight stricture.15 As a consequence, the stent is introduced through an outer sheath. The stent could be placed by either endoscopic retrograde insertion or through a nephrosotomy tract. When antegrade insertion is performed, a nephrostomy is established and a nephrostogram is performed. Using a combination of guidewires, a guidewire (hydrophilic-coated guidewire of choice) is negotiated through the stricture and passed into the bladder. A 6Fr manipulation catheter is inserted over the guidewire. The guidewire is exchanged with a superstiff guidewire and the catheter is removed. Then, a co-axial system of catheter/ sheath (an inner 6Fr ureteric catheter and outer 8.3Fr introducer sheath) is passed over the superstiff guidewire into the bladder. The guidewire and the inner catheter are removed and the Resonance stent is then introduced through the introducer sheath (outer sheath of the insertion system). Since there is no retrieval mechanism with the deployment kit, it is important to check the proximal end and to avoid pushing the ureteric stent too far.15 Once the proximal end is placed within the calyx, the final step is to remove the introducer sheath over the pusher while not moving the pusher from its position. A marked site on the introducer sheath shows the existence of only the proximal pigtail left in the sheath. Further removal

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of the sheath over the pusher allows the formation of the final pigtail in the collecting system. The retrograde approach for the insertion of a Resonance stent in the ureter is the same in concept as the antegrade technique. Radiographic and laboratory examinations immediately after the stent insertion and on a scheduled basis, as described for the conventional MSs, are also necessary for the follow-up of patients carrying Resonance stents.

5.5

Complications and problems

5.5.1 Urothelial hyperplasia Urothelial hyperplasia has proven to be a major problem following the insertion of a metal prothesis in the ureter. Ureteral patency is influenced by the hyperplastic tissue, in some cases resulting in restenosis of the stented ureter (Fig. 5.5). The phenomenon of hyperplasia has been verified endoscopically. An experimental study in pigs published by Desgrandchamps et al.25 showed that out of eight stents inserted, only one was patent 35 days after the insertion. The grade of urothelial hyperplasia seems to be influenced by the degree of force exerted on the ureteral wall as well as by the extent of the overstretching of the ureter and the subsequent urothelial trauma.26 Consequently, avoidance of overextension of the strictured area

5.5 Excretory urography: urothelial hyperplasia jeopardizing luminal patency.

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of the ureter and careful insertion is necessary. Regression of the urothelial hyperplasia 4–6 weeks after the MS implantation has been observed, thus the narrowing of the ureteral lumen is avoided.4,5,8–10 Moreover, Flueckiger et al.7 reported that reactive edema, not hyperplasia, of the urothelium during the first 2 weeks occurs and leads to constriction of the ureteral lumen. A trumpet-like configuration of the ureter adjacent to the upper extremity of the MS, which does not jeopardize ureteral patency, has also been described.10,16 Autopsy findings in a patient with malignant ureteral stricture (the patient died 2 months after implanation) demonstrated a fibrotic reaction in the lumen of the balloon-expandable stent. This phenomenon could probably be attributed to overextension of the ureteral wall resulting in urothelial trauma.4 In general, the effect of the MS on ureteral dynamic movement remains controversial. Foreign material implanted in a peristaltic organ increases the peristaltic movement of the organ in an effort to ascertain lumen patency. It has been proposed that the aperistaltic segment of the stented ureter in conjunction with the urothelial peristalsis results in urinary stasis and consequently encrustation of the uncovered areas of the MS.27

5.5.2 Encrustation Encrustation of the MS is a major problem limiting the long-term use of the ureteral endoprosthesis. The risk of encrustation is associated with both the biomaterial and the patient’s history. A comparison of the resistance of different biomaterials to encrustation by Tunney et al.55 concluded that an encrustation-resistant stent should be formed by a biomaterial resistant to bacterial infection and subsequent biofilm formation. The presence of encrustation on a small area of Wallstents that were not covered by endothelium was reported by Pauer and Lugmayr.9 The inadequate endothelial cover of the stent was caused by inefficient embracement of the ureteral wall by the MS.

5.5.3 Migration Covered MSs were introduced in an effort to overcome the issue of endothelial hyperplasia. The incidence of migration of coated MSs was observed to be as high as 81.2% in one study.16 Barbalias et al.16 implanted coated endoprostheses in 16 patients with malignant ureteral obstruction. Migration of the MS occurred in 13 of the cases causing ureteral obstruction and subsequent ipsilateral lumbar pain. The floating stents were removed cystoscopically from the bladder. Bare stents were implanted and achieved

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ureteral patency for a mean follow-up time of 8 months.16 The high and fast migration of the coated stents into the bladder, confirmation of the inadequate anchorage of the coated MSs due to the presence of the coating material and the enhanced peristalsis of the stented ureter, remain strong arguments against the use of coated stents for the management of ureteral strictures.

5.6

Virtual endoscopy and metal stents

Virtual endoscopy (VE) is a non-invasive technique that amplifies the perception of cross-sectional images, acquired by axial computed tomography (CT), in the 3-D space, providing precise spatial relationships of pathological regions and their surrounding structures. The use of appropriate software and relative algorithms is responsible for the production of virtual reality images that allow endoluminal navigation through a hollow system and simulation of conventional endoscopy. Endoluminal, as well as extraluminal, adjacent structures in all directions can be presented with VE. Body regions inaccessible to or incompatible with conventional endoscopy can also be depicted by VE. In general, VE is capable of providing information on many hollow anatomical structures and has already been used for the exploration of trachea, colon, aorta, brain ventricles, nasal cavity and paranasal sinuses.54,56–62 The application of spiral computed tomography (SCT) and MRI allows the acquisition of continuous and complete sets of raw data. The data obtained are collected by a computer workstation for post-processing and analysis. When the final 3-D dataset is available, computer algorithms generate 3-D images. The generated images contain the information obtained by either the SCT or the MRI scan. The most commonly applied techniques are shaded surface display, maximum or minimum intensity projection and volume rendering.54,56–59 VE is one of the most recent, innovative post-processing techniques and provides supplementary information to the above-mentioned diagnostic tools (SCT and MRI). The main goal of VE was the development of a noninvasive diagnostic method that would be easily tolerated by the majority of patients. The method is capable of providing images similar to those acquired by conventional endoscopy. Recent studies have reported experiences with the use of VE for the evaluation of ureteral patency after the treatment of upper urinary tract obstruction with the application of self-expandable metallic stents (Fig. 5.6).33,53 VE findings were similar to those of excretory urography. Accurate 3-D visualization of the stented area, the proximal ureter cephalad and caudal to the stent was possible by VE. Moreover, visualization of the site of interest from different angles was also feasible. Even though VE

124 (a)

Biomaterials and tissue engineering in urology (b)

5.6 Virtual endoscopy: (a) patent stented ureteral lumen; (b) stenotic distal part of the endoprosthesis.

provided information about the presence of intraluminal stenosis, the authors considered that a major disadvantage of VE was the inability of the method to differentiate in the CT acquisition settings structures with similar absorbing characteristics. In fact, ureteral wall structures were presented in the VE images with similar densities regardless of the underlying histopathology (normal urothelium, luminal encrustation, mucosal hyperplasia or tumor infiltrations). The fine detail in the epithelial lining of anatomical structures visualized with conventional endoscopic procedures cannot be differentiated and depicted by VE. However, an important advantage of VE in comparison with upper urinary tract endoscopy is the limited invasiveness. Excretory urography is probably inferior to VE. The quality of data and VE images acquired from stented ureters were superior owing to dilatation provided by the stent. The minimal mass density of the metal stents is responsible for the reduced scattering of the X-rays’ quantum energies. Specially modified CT reconstruction protocols are able to overcome artifacts from strut reflections.63 It is important to note that the application of VE in the urinary tract involves certain limitations. Small and flat lesions or mucosal thickening are depicted with difficulty, while biopsy tissue specimens for histopathological examination cannot be provided. Moreover, the bladder must be sufficiently dilated. Analysis of both the axial and virtual images is usually necessary for optimal evaluation. The risk of radiation is always present with the CT-based VE, only MRI can overcome this problem. Conditions that hinder the excretion of the contrast material into the upper urinary tract, such as renal insufficiency or high-grade tumor obstruction, do not allow the acquisition of VE images. Nevertheless, if endoluminal stenosis

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or obstruction is present, virtual endoluminal navigation both cephalad and caudal to the stenotic segment is possible through VE.

5.7

Extra-urinary drainage of the upper urinary tract

Metal stents represent an alternative solution for percutaneous nephrostomy in difficult cases that cannot be managed by double-J catheter insertion or in cases requiring long-term urinary drainage due to ureteral obstruction. Nevertheless, another method has also been introduced for the permanent management of both malignant and benign cases of ureteral obstruction. The extra-anatomic subcutaneous stents (EASSs) have been used in a limited number of cases.64–69 Subcutaneous urinary diversion with the extra-anatomic stents has been performed in an effort to achieve longterm urinary drainage from renal pelvis to bladder.64,65 Indications for the insertion of extra-anatomic stents are cases of impassable ureteric obstructions or complete disruption of the ureter. Two types of extra-anatomic stents have been introduced: temporary and permanent.66 The temporary stent used is the 8Fr Paterson Forester stent (Cook Ireland, Limerick, Ireland) which is a 65 cm long polymeric stent with a double-J design without side holes. Regular replacement of the stent is necessary with intervals ranging between 6 and 12 months. The stent has proven to be susceptible to encrustation and lumen obstruction.66 The permanent stent used is the Detour stent (Mentor Porgès, Coloplast, Lancing, West Sussex, UK) and this has been used for both benign and malignant disease that could not be otherwise treated.67–70 The Detour stent is composed of an outer 27Fr polytetrafluoroethylene (PTFE) silicone tube and an inner 17Fr silicone tube that extends in both ends beyond the PTFE tube. For insertion of the stent, a nephrostomy is established and the stent is inserted in the renal pelvis with the use of sheaths. The development of a subcutaneous tunnel follows and the stent is introduced into the tunnel. A Pfannenstiel incision is made and the stent is inserted in the bladder through a cystotomy.71 Lloyd et al.71 recently published their experience with the Detour stent.71 They used the stent in eight patients. One patient had Detour stents inserted bilaterally. The etiology of the ureteric obstruction was persistent malignant disease in three patients and complicated benign disease in five patients. Patients with malignancy had metastatic breast cancer, retroperitoneal sarcoma and carcinoid tumor of the colon complicated with fibrosis of the ureter. No complication was observed in any of these patients, the former two died with the stent in place while the latter was still alive at the end of the follow-up period. Patients with benign disease had intraoperative ureteral injury in three cases, retroperitoneal fibrosis due to previous MS insertion and persistent UPJ obstruction (previous open and endoscopic surgery)

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in a single kidney and were treated with the Detour stent. One of the patients with benign disease was diagnosed with high-grade transitional cell cancer of urothelial origin and prostatic cancer; he died during the followup period. Complications observed in this series were hematuria and infection. Experience with Detour stents inserted for palliative management of malignancies has also been reported by Deshgrandchamps et al.72 A total of 27 stents were inserted in 19 patients. Insertion of the stent was successful in all cases while no intraoperative and immediate postoperative complications were observed. Mean follow-up time was 7.8 months. Only four patients were alive 1 year after stent insertion, the remaining patients died as a result of their malignancies. Suprapubic parietal infection was observed in patients with bladders affected by tumor or previous irradiation. Quality of life was reported to be improved. A composite prosthesis consisting of two co-axial tubes – an inner pure smooth silicone tube and an outer coiled PTFE tube – designed to serve as a ureteral substitute was inserted for the treatment of 22 patients with malignancies and 5 with benign diseases.73 Stent insertion was performed percutaneously through the renal pelvis, a subcutaneous tunnel was formed and finally the stent was introduced into the bladder by suprapubic incision. Mean follow-up period for patients who died was 6.3 months; the remaining patients had a mean follow-up time of 47 months. Complications observed were removal of stent due to skin erosion and return to percutaneous nephrostomy in two patients owing to tumor progression. Patients with long-term follow-up did not suffer any complications.

5.8

Future trends

Following the first implantations of MSs in the field of urology, significant progress has been made in improving stent design in an effort to minimize complications and problems of the application of MSs in a delicate organ such as the ureter. Every attempt could be considered as a step towards the development of the ideal MS. The use of coated or covered MSs and the introduction of bioabsorbable stents seems promising in minimizing the morbidity associated with stent implantation while improving urinary drainage into the bladder.47,48,74–76 Pharmacologically active agents can be incorporated on the stent surface, either directly on the polymeric surface of the stent (coated MSs) or into the core of the polymeric structure of the stent (DES). In the first case, the pharmacological agent is capable of preventing infection or encrustation, whereas in the second case, the pharmacological agent is delivered in a sustained-release fashion and may affect the urothelium proliferation.

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Recent experimental studies have focused on the field of DESs. The application of DESs in the pig ureter and the canine urethra is a glance into the future. DESs are already included in the clinical practice of interventional cardiology and clinical evaluation in urology should be expected. The interesting capability of DESs to reduce neo-intimal hyperplasia and the promising results in animals in reducing urethral and ureteral tissue hyperplasia, as reported by Shin et al.48 and Liatsikos et al.,47 represent an important challenge for the clinical application of DESs. Antimisiaris et al.77 have proposed a method of controlled release of drugs by liposome-covered MSs.77 The pharmacological agent is loaded on the liposome particles and the liposomes are applied on the surface of the stent. Controlled release of the lipid and drug components of the liposome particles was achieved, rendering the stent-associated liposomal drug formulations as slow-release depots, an efficient method for treating complications of ureteral stent obstruction. Interesting methods for the coating of MSs with liposomes and preparing them as DESs have also been investigated by the same group.78,79 In general, the DESs could probably solve many of the problems of stenting in the urinary tract, but extensive animal and human trials are needed to prove their efficacy. In addition, the introduction of retrievable stent designs is needed for easier removal of DESs.48 Recently, the development of a tissue-engineered, balloon-expandable, autologous tissue-covered stent has been reported.80 In vivo architecture technology and covered stent technology was used for the development of the stent. Tissue covering probably enhances early vascular reconstruction in the stent lumen which is responsible for the reduction of stent restenosis frequency. Further investigation is already being conducted. A summary of the existing experimental studies is presented in Table 5.3. Table 5.3 Experimental studies in ureteral stenting Investigator group Wright et al.24

Desgrandchamps et al.25

Issue

Description

Implantation of self- and balloon-expandable metal stents (MSs) in normal/stenotic canine ureters Implantation of selfexpanding MSs in pig ureters

Complete occlusion in all stented ureters due to mucosal hyperplasia in 4–8 weeks

Distal narrowing was observed owing to functional discrepancy between the adynamic stented ureter and the normal underlying ureter; urothelial hyperplasia was also present without playing any important role in obstructing the ureter

Table 5.3 Continued Investigator group Liatsikos et al.

35

Thijssen et al.26

Antimitsiaris et al.77

Leveille et al.36

Liatsikos et al.48

Koromila et al.78

Antimitsiaris et al.79

Blaschko et al.50

Nakayama et al.80

Issue

Description

Implantation of bare MS and coated MS (internally and externally) in normal pig ureter Implantation of selfexpanding MSs in histologically normal canine ureters

Coated stents result in minimal tissue ingrowth; tendency to migrate towards the bladder; bare MSs cause less inflammation of surrounding tissues Ureters remained patent; MS was not incorporated within the ureter wall; penetration of epithelium/ submucosa between wire struts and areas of fibrosis in submucosal layer were observed A possible efficient method of treating ureteral stent obstruction

Controlled release of dexamethasone after the implantation of liposome-covered MS while simulating in vivo conditions Bare MS and MS lined with porous biocompatible polymer in canine model Implantation of bare MS and paclitaxel drug-eluting stent (DES) in normal pig ureter Coating polymer-covered stents with heparinencapsulating liposomes Controlled heparin delivery Evaluation of the efficiency of liposome vesicles in the coating of polymer-coated stents, controlled drug release, hemocompatibility of the stents Comparative evaluation of the urinary flow of Resonance stent and standard ureteral stent

Development of in vivo tissue-engineered, autologous tissuecovered stents (biocovered stents)

MSs lined with a biocompatible material help prevent tissue ingrowth; in contrast, bare MSs cause ureteral obstruction and hydronephrosis Paclitaxel DES generated less inflammation and hyperplasia of surrounding tissues resulting in maintenance of ureteral patency Controlled released of heparin is achieved via multilamellar vesicles or dehydration–rehydration vesicles; the latter method may be functional under high flow conditions Liposomal formulations of drugs can be used as coating in polymercovered stents; the release rate can be modified through the modification of liposome characteristics; heparinencapsulating liposomes improve hemocompatibility markedly The Resonance stent provided less flow than the standard stent; nevertheless, under circumstances of extrinsic ureteral compression sufficient to acclude the standard stent, the Resonance continues to procide satisfactory drainage It was considered that normal vascular reconstruction may be enhanced with high reliability; as a consequence the rate of intraluminal restenosis would be reduced

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In conclusion, MSs in the upper urinary tract are certainly a useful tool for the modern urologist. However, the fate of MSs in the future seems to be strongly related to the development of new designs and materials that can improve the outcome and expand the uses of MSs.

5.9

References

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73 jabbour me, desgrandchamps f, angelescu e, et al. Percutaneous implantation of subcutaneous prosthetic ureters: long-term outcome. J Endourol 2001;15(6):611–4. 74 laaksovirta s, talja m, valimaa t, et al. Expansion and bioabsorption of the self-reinforced lactic and glycolic acid copolymer prostatic spiral stent. J Urol 2001;166:919–22. 75 lumiaho j, heino a, pietlainen t, et al. Themorphological in situ effects of a self-reinforced bioabsorbable polyactide (SR-PLA 96) ureteric stent: an experimental study. J Urol 2000;164:1360–3. 76 lumiaho j, heino a, tunninen v, et al. New bioabsorbable polyactide ureteral stent in the treatment of ureteral lesions: an experimental study. J Endourol 1999;13:107–12. 77 antimitsiaris sg, siamblis d, liatsikos e, et al. Liposome-coated metal stents: an vitro evaluation of controlled-released modality in the ureter. J Endourol 2000;14:743–7. 78 koromila g, michanetzis gpa, missirlis yf, et al. Heparin incorporating liposomes as a delivery system of heparin from PET-covered metallic stents: effect on haemocompatibility. Biomaterials 2006;27:2525–33. 79 antimitsiaris sg, koromila g, michanetzis g, missirlis yf. Liposome coated stents: A method to deliver drugs to the site of action and improve stent bloodcompatibility. J Liposome Res 2006;16:303–9. 80 nakayama y, zhou ym, ishibashi-ueda h. Development of in vivo tissueengineered autologous tissue-covered stents (biocovered stents). J Artif Organs 2007;10:171–6.

6 Coated ureteral stents F. C AU DA, Ospedale Koelliker, Italy; V. C AU DA, LudwigMaximilians-Universität München, Germany; and C. F I O R I, Azienda Ospedaliero-Universitaria San Luigi Gonzaga, Italy

Abstract: In this chapter we discuss the role of heparin and diamond-like amorphous carbon coatings in reducing ureteral double-J stent encrustations. Morphological, compositional and qualitative analyses have been carried out using field emission scanning electron microscopy, energy dispersive spectroscopy and micro-infrared spectroscopy, both before stent insertion and after 1 month of indwelling. Key words: ureteral stent, heparin coating, diamond-like amorphous carbon coating, scanning electron microscopy, infrared spectroscopy, stent encrustation.

6.1

Introduction

Ureteral stenting for urinary drainage has become a routine procedure in urology. In general, it is used as a tool to protect the ureter: a stent is not used as a therapy, but as an auxiliary device in urology and endourology to manage patient urinary disease better (Arshad et al. 2006). Therefore clinical indications for the use of a stent may be wide-ranging, from urolithiasis to reconstruction, trauma and transplantation. However, all these clinical indications have the common aim of improving urine drainage in the following clinical cases: •

• •

acute ureteral obstruction (i.e. in a stone-former patient until the obstructing stone is removed, or one with a ureteral clot in the case of hematuria etc.); permanent ureteral obstruction (extrinsic or intrinsic); post-surgical treatment (i.e. ureteroscopy, ureteral surgery, etc.).

Obviously, when a lengthy time of indwelling is required, i.e. 1 month or even more, encrustation and occlusion of the catheter lumen can occur. Stent encrustation is the major problem and an extremely difficult one to prevent, therefore the prevention of bacterial adherence to biomaterial surfaces is an important topic. 134

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The most important step in the encrustation of any urinary drainage device seems to be the formation of biofilm, which is a complex aggregation of micro-organisms growing on a solid substrate with their extracellular products accumulating on the implant surface (Costerton 2007). Biofilms are characterized by structural heterogeneity, genetic diversity, complex community interactions and an extracellular matrix of polymeric substances. This matrix protects the cells within it and facilitates communication among them through chemical and physical signals. Bacteria living in a biofilm can have significantly different properties from free-floating bacteria, both in behavior and in phenotypic form, as the dense and protected environment of the film allows them to cooperate and interact in various ways. This leads to a reduced susceptibility to prophylactic antibiotics and an increased resistance to detergents, as the dense extracellular matrix and the outer layer of cells protect the interior of the community (Tenke et al. 2004). The first stage in biofilm formation is bacterial colonization: initially micro-organisms are deposited on the biomaterial surface and then adhere to it through weak, reversible Van der Waals forces from exopolymer production. They then grow, multiply and colonize the entire surface, giving rise to the bacterial biofilm. The biomaterial surface characteristics, the presenting clinical condition, the bacterial interface features and their behavior all influence bacterial adhesion to the surface of a foreign material. The second stage after biofilm formation is the production of ammonia by urease-producing bacteria, infecting the urinary tract and the stent. The consequent pH increase of the urine causes the precipitation of inorganic compounds, such as the crystallization of calcium oxalate, magnesium ammonium phosphate (struvite) and calcium phosphates (hydroxyapatite and brushite). The development of these encrustations can obstruct the stent and impair the urinary flow, which compromises patient care and may lead to kidney infections, sepsis and shock (Warren et al. 1994). Temporary solutions that can be used to prevent stent encrustation include exchanging the stent at regular intervals, modifying the type or size of catheters, increasing fluid intake, administering acidifying drugs and washing the catheter and bladder with acidic, antiseptic or saline solutions (Arshad et al. 2006). Administering antibiotics at the time of stent insertion is one method of preventing the complication of infection. It has been reported that the antibiotics adhere to the stent surface at bactericidal levels and can delay the adherence of bacteria to stents, but their use requires patient agreement and may cause morbidity (Reid et al. 2001). Since all of these approaches are mostly ineffective, stent surface modifications to prevent bacterial and inorganic molecule adhesion have been developed. Various strategies have been set up, such as silver-coated surfaces (Leung et al. 1992; Imram 1999), the locally controlled release of

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antibiotics or antiseptics (Raad et al. 1997; Cormio et al. 2001; Leung et al. 2001; John et al. 2007), and surface modification to change hydrophobicity (Jansen et al. 1993) or to create functional groups with intrinsic antimicrobial activity. As far as antibiotics are concerned, novel biomimetic and bioactive silicones are under development to allow the incorporation of antibiotics (Gorman and Woolfson 2002). DiTizio et al. (1998) described a liposomal hydrogel-coated biomaterial, the liposomes of which contain ciprofloxacin, and this has been shown in vitro to resist encrustation. One group has produced a gentamicin-releasing urethral catheter, which has been shown in a rabbit model to inhibit encrustation in the short term (Cho et al. 2001). The use of silver as a means of preventing bacterial colonization and encrustation has been the subject of much work. Silver alloys can be coated on to the surface of biomaterials. One group observed the effect of a silverreleasing device in prostheses within a dynamic model, using human urine. It found that the spread of microbes along the surface of the biomaterial was prevented, in the short term, from passing beyond the device (Stickler et al. 1996; Stickler 2000). This finding has been confirmed by other studies (Cormio et al. 1996). However, Sabbuba et al. (2002), in contrast, suggested that the swarming of Proteus over Foley catheters is not inhibited by silver. In addition, Cormio et al. (1996), using an in vitro static human urine model, have shown that coatings containing silver do not impair bacterial adhesion to biomaterial. This controversy prompted one meta-analysis of the effectiveness of silver-bearing catheters. Saint et al. (1998) demonstrated clearly that silver alloy-coated urethral catheters are effective in decreasing the rate of urinary tract infection in patients with long-term catheters, with a consequent decrease in the occurrence of complications. However, this has to be weighed against the slightly higher cost of these prostheses. Polymers that have been used for hydrogel coatings include polyacrylamide, polyvinyl alcohol, polyethylene glycol and poly-N-vinyl pyrrolidone (Watterson et al. 2002). The trapping of water within the polymeric structure results in a decrease in the coefficient friction of the biomaterial, which improves biocompatibility by reducing cell and biomaterial–tissue interface adhesion. Nevertheless, Choong et al. (2000) noted that a hydrogel-coated stent accumulated significantly more encrustation that a traditional stent. The use of heparin-coated surfaces is a very effective strategy for preventing bacterial anchorage (Appelgren et al. 1996; Riedl et al. 2002; Cauda et al. 2008). Heparin is an anti-thrombogenic agent with strong electronegativity that prevents cell adhesion. Such heparinization can be a practical and economical approach to the prevention of catheter-associated bacteremia or fungemia. Plasma-deposited diamond-like amorphous carbon coatings are well known for their excellent biocompatibility; a recent study reported on the ability of diamond-like carbon to decrease the formation of crystal-

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line bacterial biofilm as well as stent-related side-effects and discomfort (Laube et al. 2006; Laube et al. 2007). In this chapter, our experience of double-J polyurethane ureteral catheters is reviewed: the aim is to evaluate in vivo the efficacy of covalently bonded heparin-coated and amorphous carbon-coated stents in reducing microbial adherence and colonization (Cauda et al. 2008). Field emission scanning electron microscopy (FESEM), energy dispersive spectroscopy (EDS) and infrared spectroscopy coupled with an optical microscope were used to characterize biofilm morphology and make a compositional analysis of the stent surface after indwelling. All results were compared with coated and uncoated stents before insertion in the ureter. A comparison of the behavior in preventing biofilm adhesion between heparin-coated stents and carbon-coated ones will also be attempted.

6.2

Methods

The study began in January 2006 at Dipartimento di Nefrourologia, ‘S.S.C.V.D. per il trattamento integrato della calcolosi urinaria’ (Stone Center) – Ospedale Maggiore S. Giovanni Battista, Turin, Italy. We enrolled 16 patients (45–73 years old): 10 showed bilateral ureteral obstruction, thus requiring bilateral ureteral stenting for 1 month; 6 patients showed unilateral obstruction and were unilaterally treated with coated stents for different periods of time. In detail: (a)

5 patients with bilateral obstruction were treated with a heparincoated stent in one side and an uncoated stent in the other side, with an indwelling time of 1 month; (b) 1 patient with chronic unilateral obstruction was treated with a heparin-coated stent in position for 10 months; (c) 1 patient with chronic unilateral obstruction was treated with a heparin-coated stent in position for 12 months; (d) 5 patients with bilateral obstruction were treated with a carbon-coated stent in one side and an uncoated stent in the other side, with an indwelling time of 1 month; (e) 4 patients with unilateral obstruction were treated with carbon-coated stents positioned unilaterally for 1 month. Each of the patients with bilateral obstruction randomly received a coated stent (with heparin, group (a); with amorphous carbon, group (d)) and an uncoated one; in this way every patient received a coated stent in one ureter and an uncoated stent in the other. Patients with a unilateral obstruction received a heparin-coated stent (groups (b) or (c)) or an amorphous carbon-coated stent (group (e)); 26 stents were therefore considered for the analyses.

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The main reasons for inserting a stent were post-endoscopic stone treatment and uretero-pelvic junction (UPJ) obstruction (awaiting surgery). In two cases (groups (b) and (c)), long-term indwelling stenting was performed for hydronephrosis due to extrinsic compression in patients with multicystic disease (no surgical indications) and UPJ obstruction after failure of endoscopic and surgical treatment (patient refused re-intervention). In all cases, double-J stents were inserted during cystoscopy. A retrograde ureteropyelography was performed at the start of procedure, to evaluate the excretory system. Ciprofloxacin (500 mg twice a day) was used for prophylaxis at day 0, +1, +2 and +3 after the procedure. In all cases, stents were removed after 1 month during cystoscopy, except in two cases (groups (b) and (c)), where the stent was removed after 10 and 12 months respectively. A record was kept of the occurrence of technical problems, violations of aseptic technique during procedures, if any, antibiotic therapy administration, other therapeutic interventions administered during the period of indwelling and the presence of fever, infections or urinary symptoms. Urine analysis and culture were performed on day +15 after the procedure. After removal, different segments of the stent were randomly cut in small sections, both longitudinally and transversally, for characterization studies. In the longitudinal sections, both the inner and outer surfaces were analyzed, while in the transversal sections, one cut surface per section was studied. The two edges of the stent were studied too, as transversal sections. Different analyses were performed: morphological, compositional and qualitative. In particular, morphological and compositional analyses were carried out on both kinds of section using FESEM (Assing FESEM Supra 25, Gottingen, Germany) and an EDS device (INCA X-Sight, Oxford Instrument, Gottingen, Germany). For these measurements, the stents had to be covered by a thin layer of gold or silver to become conductive. The thickness of stent encrustation was measured on transversal sections by INCA® electronic imaging software during the FESEM photo session; if encrustation was present but not measurable, the thickness was considered as zero. As the encrustation thickness was variable, measurements were carried out at different points, taking the highest thickness value registered per transversal section for the statistical analysis. For each longitudinal section, encrustation was measured as a percentage of coverage of the stent surface using PC software for image analyses. Statistical analyses to evaluate the differences between the two groups (coated versus uncoated stents) were carried out using Student’s t test, the chi-square test and Fisher’s exact test. All data were analyzed using a PC application (Statistica, Microsoft, Tusla, OK) run on a PC. Compositional and qualitative studies were carried out with a micro-infrared (IR) spectrophotometer (Bruker Optik, Ettingen, Germany) in attenuated total reflectance (ATR) mode. This instrument first allows investigation of the sample

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surface cut longitudinally with an optical microscope and then analysis of its composition through the IR spectroscopy: this allows mapping of both the inner and outer surfaces of the stents, so completing the information derived from FESEM and EDS. For comparison, the same techniques were applied to coated stents before insertion.

6.3

Results

6.3.1 Clinical and operative data No technical problems or violations of asepsis were recorded during endoscopic procedures. In all patients belonging to groups (a) and (d), retrograde ureteropyelography performed at the start of the procedure showed bilateral ureteral obstruction with various degrees of excretory system dilatation. During the study period, no patient reported fever or flank pain. Urine culture was negative in all cases. In one patient in group (a) and two patients in group (e), frequency/urgency symptoms were recorded. The patients were successfully treated with anti-muscarinics until the removal of the stents. Stents were removed without technical difficulties in all cases. No technical problems occurred during the endoscopic procedures on the two patients with chronic unilateral obstruction (groups (b) and (c)). In these cases, ureteropyelography showed ureteral obstruction with severe excretory system dilatation. During the study period, no fever, flank pain or urinary symptoms were recorded and urine culture was negative in both cases.

6.3.2 Heparin-coated stent analysis Before insertion Polyurethane heparin-coated stents were morphologically characterized by SEM and elemental analysis using EDS. They showed a quite regular surface, with some irregular particles (polyurethane or impurities) of about 0.5–1 μm. The heparin layer could be observed in a cross-section of the stent on the polyurethane substrate: its thickness was estimated at 5 μm, while pores of about 400–500 nm were present on the heparin surface (see Fig. 6.1). After 1 month of indwelling Five stents were bilaterally implanted for 1 month, together with uncoated polyurethane stents, in order to evaluate the ability of heparin to prevent biofilm and inorganic encrustation deposition. After removal, the heparincoated stents were morphologically, compositionally and qualitatively analyzed. Figures 6.2 and 6.3 show different heparin-coated stents after 1 month

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200 nm

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Date:25 Jan 2006 Signal A = InLens Time:16:17:17 Stage at T = 45.0° User Name = ANGELICA

6.1 Electron microscopy characterization of heparin-coated polyurethane stent before indwelling. (a) Inner surface; (b) magnification of the porous texture of the heparin layer.

of indwelling in the patients. The levels of surface encrustation really depend on patient pathology, but, in general, the internal surfaces were more encrusted than the external ones. In the FESEM analysis of the transversal sections, encrustation varied from zero, where the heparin coating was still distinguishable at the cut edge of the stent, to 8.5 (±0.6) μm. The average encrustation extension of the longitudinal section surface was 67% (±33%), measured on both inner and outer surfaces. The elemental EDS analysis confirmed the presence of the heparin layer at the cut edges of the stent. In one case, oxygenated compounds of calcium were clearly identified on the inner surface of a heparin-coated stent. EDS analysis enabled this compound to be identified as calcium oxalate, for which the chemical formula is CaC2O4: the amount of elements as an atomic percentage clearly reflects the theoretical amount (Table 6.1) in the chemical formula. The micro-IR spectra of heparin-coated stents after insertion were obtained for both the internal and external longitudinal sections and compared with the spectrum of a stent before insertion. An example is showed in Fig. 6.4. All the spectra of the coated stents collected in a non-encrusted

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(b)

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6.2 Two transversal cross-sections of the heparin-coated stents, showing a level of encrustation.

zone (spectrum (b)) were very similar to the spectrum of the heparincoated stent before insertion (spectrum (a)), meaning that no alteration of the heparin surface occurred during 1 month in urine media. In the spectral zone from 1200 to 600 cm−1, vibrations of chains of single bonds are found: the bands in this range are typical of a polymer chain, in this case of polyurethane and heparin. In contrast, the spectrum collected from an encrusted zone of the inner surface (spectrum (c)) was different from that of the heparin-coated stent before insertion. The presence of an organic bacteria

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(a)

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Mag = 82 X Signal A = InLens Date: 14 Jul 2006 Time: 10:40:10

6.3 Internal surface of the longitudinally cut heparin-coated stents. (a) A quite encrusted surface with (b) magnification of the inorganic deposits; (c) low levels of biofilm formation and (d) magnification at the edge of the cut surface to measure biofilm thickness.

Table 6.1 Ratio of elements in the chemical formula of calcium oxalate and from data obtained by EDS analysis for a heparin-coated stent after 1 month of indwelling Element

Theoretical atomic %

Experimental atomic %

O C Ca

57.1 29.1 14.3

63.92 21.55 14.53

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C— — O C—O

Absorbance (a.u.)

(e)

—CH2, —CH3 (d) H2O

(c) (b) (a) 3500

3000

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6.4 Micro-IR spectra collected at the longitudinally cut heparin-coated stent surface. (a) Surface of the stent before insertion; (b) internal surface of the stent after 1 month of indwelling in a non-encrusted zone; (c) biofilm-encrusted zone of the internal surface; (d) external surface of the heparin-coated stent; (e) calcium oxalate deposition on an internal surface. a.u., arbitrary units.

biofilm was evidenced by the increase of the water content at 3280 cm−1 and by the peaks at 3000–2800 cm−1, belonging to the stretching vibrations of —CH3 and —CH2 groups (see the dashed rectangle). These peaks are generally related to organic material, which can be either polyurethane or bacteria biofilm. In the present case, the increase of —CH3 and —CH2 stretching vibrations from spectrum (a) to spectrum (c) is related to the deposition of an organic film on to the polyurethane stent surface, thus forming the biofilm. Moreover, peaks of heparin and polyurethane are no longer recognizable in the zone under 1200 cm−1 (called the ‘fingerprint’ zone, in which bands typical of the polymer chain are observed), therefore the encrustation depth is more than 1 μm – where 1 μm is the maximum depth of IR in the ATR mode used here. The external surface presented, in general, very similar spectra (see spectrum (d)) to that of the stent before insertion, confirming the findings of electron microscopy characterization. In only one case was the presence of calcium oxalate (spectrum (e)) identified on the internal surface of a heparin-coated stent. The spectrum of this inorganic compound is very specific and therefore easy to recognize. Typical vibrations of C=O and

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C—O groups, present in the molecule, are observed at 1613 and 1310 cm−1 respectively. As no peaks of heparin and polyurethane were recognizable in this spectrum, the calcium oxalate was thicker than 1 μm. After prolonged indwelling (10 and 12 months) Two patients, who needed stenting for a longer period of time than 1 month, were studied and their heparin-coated stents were characterized after 10 and 12 months of indwelling respectively. Neither of the coated stents showed any kind of encrustation or degradation of the heparin surface. The cross-sections of the catheters were visibly free of biofilm, as were the drainage pores on the outer surface (Fig. 6.5). Elemental analysis was carried out on the internal surface, which is usually the most encrusted part of the stent, and in both cases showed only the presence of carbon and oxygen, the components of polyurethane (Fig. 6.5(c)). In contrast, in the case of the stent after 12 months of indwelling, the outer surface was rich with bacteria biofilm accompanied by sodium chloride and oxygenated compounds of calcium, as shown by the EDS spectrum in the inset of Fig. 6.5(d). Micro-IR spectra confirmed the data obtained from electron microscopy and elemental analysis, showing, moreover, that the peaks were very similar to those in the spectrum before insertion. We concluded that no significant modification of the heparin layer surface occurred after 10 and 12 months of insertion (Fig. 6.6, spectra (b) and (c)). The IR spectrum of the outer surface of the 12-month stent showed the typical peaks of bacterial biofilm (spectrum (d)), as already observed by morphological analysis. The increased amount of water in the range between 3400 and 3100 cm−1, the decrease of —CH2 and —CH3 stretching vibrations between 3000 and 2800 cm−1, and the disappearing of the typical vibrations of polyurethane and heparin in the ‘finger print’ zone below 1200 cm−1 mean that an encrustation layer was deposited on to the stent surface.

6.3.3 Carbon-coated stent analysis Before insertion Both the internal and external surfaces of a carbon-coated stent were characterized by means of FESEM and EDS analyses. Figure 6.7 shows the external surface, near to a lateral drainage hole, and the internal surface of the stent. At higher magnifications, the inner surface showed several grains belonging to polyurethane (see the inset of Fig. 6.7(b)). Although the polyurethane surface seems to be quite flat macroscopically, at higher magnifications a rougher surface is shown.

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6.5 Scanning electron micrographs of the two stents after prolonged indwelling. (a) Crosssection of the 10-month stent; (b) lateral drainage hole of the 10-month stent; (c) inner surface of the 12-month stent with elemental analysis in inset; (d) outer surface showing the biofilm encrustation on the 12-month stent with EDS spectrum in inset.

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—O C—

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6.6 Micro-IR spectra of stents after prolonged indwelling. (a) Surface of the stent before insertion; (b) internal surface of the stent after 10 months of indwelling; (c) internal surface of the stent after 12 months of indwelling; (d) external surface of the 12-month stent with biofilm encrustation.

After 1 month of indwelling Of the nine carbon-coated stents analyzed after 1 month of indwelling, four of them showed remarkable levels of encrustation, due to particular patient pathology, while the other five were completely clean. Figure 6.8 presents some examples of stents free of encrustation, showing the cross-section and part of the internal and external surfaces of the stents. The elemental analysis carried out on the internal surface (see inset of Fig. 6.8(d)) showed large amounts of carbon and oxygen, derived from the polyurethane stent surface, and sodium chloride (NaCl). Figure 6.9 shows carbon-coated stents with considerable levels of encrustation after 1 month of indwelling. The encrustation is, in general, not as compact and uniform as that found in the bilateral untreated stent. This finding can be easily explained by assuming that the presence of the carbon treatment on the stent surface prevents encrustation and bacterial adhesion. Nevertheless, a certain degree of encrustation is present. Encrustation thickness is between 10 and 38 μm (see Fig. 6.9(b) and (d)). As shown by the elemental analysis in Fig. 6.9(g), the encrustation is mostly made up of different salts, deposited from urine. Even when there were a lot of elements present in this encrustation, in some

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6.7 Morphological characterization on carbon-coated stent before indwelling; (a) external surface in proximity of a drainage hole; (b) inner surface and (c) its magnification.

cases it was possible to detect the presence of brushite from the table of atomic percentages. Relative ratios between calcium, phosphorus and oxygen are reported in Table 6.2 and compared with the ratios obtained from the chemical formula of brushite (CaHPO4.H2O). In Fig. 6.9(c) it is possible to observe the high level of encrustation of the external surface of the encrusted stent. In another case, as seen in Fig. 6.9(e) and (g), the internal surface shows varying degrees and forms of crystalline deposits. The EDS spectrum reported in Fig. 6.9(h) shows the presence of oxygenated compounds of calcium. The atomic percentages do not reflect

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6.8 FESEM on the carbon-coated stents that were free of encrustation. (a) and (b) Crosssections; (c) outer surface in proximity of a lateral drainage hole; (d) inner surface with elemental analysis in inset.

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6.9 Micrographs related to the encrusted carbon-coated stents. (a) Cross-section with (b) magnification to measure the encrustation thickness; (c) outer surface with (d) lateral drainage hole magnification; (e) inner surface with (f) magnification of calcium oxalate crystals; (g) elemental analysis on the encrustation depicted in panel (c); (h) EDS analysis on the encrustation of panel (f).

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Table 6.2 Ratio of elements in the chemical formula of brushite and from data obtained by the EDS analysis from the micrograph shown in Fig. 6.9(c)

Element ratio

Chemical formula

From EDS analysis (Fig. 6.9(c))

Ca/P Ca/O P/O

1 0.17 0.17

1.17 0.11 0.09

Table 6.3 Ratio of elements in the chemical formula of calcium oxalate and from data obtained by EDS analysis for an encrusted carbon-coated stent as depicted in Fig. 6.9(f) Element

Theoretical atomic %

Experimental atomic %

O C Ca

57.1 29.1 14.3

56.72 34.10 8.70

exactly the amount of the calcium element in the calcium oxalate compound (CaC2O4), which is theoretically 14.3%. The percentages of the other two elements, carbon and oxygen, are, in contrast, nearer to the theoretical values, as reported in Table 6.3. In addition to calcium oxalate, sodium chloride was also precipitated. Micro-IR analysis was performed on the carbon-coated stents after 1 month of indwelling. The spectra obtained were compared with the spectrum of the stent before use, in order to find any modifications of the polyurethane stent surface and identify bacteria biofilm and inorganic compounds deposited from urine. Figure 6.10 shows a spectrum collected in a non-encrusted zone (spectrum (b)), which is very similar to the spectrum of the carbon-coated surface before insertion (spectrum (a)). The same can be seen for the spectrum of the external stent surface (spectrum (c)), which is quite free of encrustation. The encrusted zones were studied using the optical microscope associated with IR spectroscopy. Spectrum (d), as an example, is quite different from that obtained from the carbon-coated stent before insertion. A broad band, from 3400 to 3200 cm−1, of hydration water is present, and —CH2 and —CH3 stretching and bending peaks are recognizable, as indicated by the dashed rectangle, proving the presence of an organic encrustation of bacteria in the zone analyzed. This spectrum did not show any peak of polyurethane, therefore the biofilm thickness on both

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—O C—O C—

(e)

Absorbance (a.u.)

—CH2, —CH3 H2O

(d) (c)

(b)

–NH

(a) 3500

3000

1500

2500 2000

1000

–1)

Wavelength (cm

6.10 Micro-IR spectra of carbon-coated stents. (a) Surface of the stent before insertion; (b) non-encrusted internal surface of the stent; (c) non-encrusted external surface; (d) inner surface with biofilm encrustation; (e) calcium oxalate spectrum.

surfaces is thicker than 1 μm, the depth of the analysis in ATR mode. Spectrum (e) was obtained from a heavily encrusted zone of the internal surface. It shows all peaks ascribable to calcium oxalate, whose presence had already been proven by means of elemental analysis with EDS. The IR spectrum of the calcium oxalate crystals appears very well defined, showing the high purity of the isolated biomineral on the internal stent surface.

6.4

Discussion

Infection, encrustation and obstruction of long-term indwelling catheters, ureteral stents and nephrostomy tubes are significant medical and economic problems. Because of these problems, about 50% of patients with permanent urinary drainage devices present as emergency cases during the course of their disease and produce costs of up to US$1.5 billion per year in Western Europe and a similar amount in North America (Kohler-Ockmore and Feneley 1996). Many anti-adherence strategies have been developed but none of the investigated strategies proved to be optimal: silver coatings, antibiotic impregnation, hydrogel coatings (Riley et al. 1995; Maki et al. 1998; Chew et al. 2006).

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In previous in vitro and in vivo studies, the heparinization of medical devices showed a reduction in microbial colonization (Ruggieri et al. 1987; Appelgren et al. 1996; Hildebrandt et al. 1999). There are many hypotheses concerning the role of heparin in reducing bacterial adhesion; in an area with low protein content, such as the urinary tract, the effect of heparin could depend on its hydrophobicity and negative charge. Ruggieri et al. (1987) showed a 90% reduction in bacterial adhesion on heparin-coated catheter surfaces, while Hildebrandt et al. (1999) demonstrated a reduction of stent encrustation by using a heparin coating in an experimental study. Riedl et al. (2002) published an interesting in vivo study, analyzing heparincoated double-J and nephrostomy tubes with electron microscopy after 2–6 weeks. The authors demonstrated that heparin-coated stents were unaffected by encrustations after 6 weeks of indwelling time, but all uncoated stents (used as controls) showed various forms and degrees of deposit, so they concluded that heparin reduces stent encrustation. Of currently known biomaterials, diamond-like carbon (DLC) a-C:H stands out owing to the variability in its composition, and its remarkable physical and chemical properties in the fields of microbiology and medicine. Various investigations have proved its excellent compatibility with respect to blood and bone cells. DLC is a thermodynamically meta-stable state of carbon, in which diamond-like (sp3-hybridized) and graphite-like (sp2hybridized) bondings coexist with a large number of sp3 bonds. DLC films can be prepared by miscellaneous deposition methods – e.g. ion deposition, sputtering, pulsed laser deposition and plasma-enhanced chemical vapor deposition (which uses accelerated hydrocarbon ions (plasma) as filmforming particles). In general, DLC films are characterized by high mechanical hardness and chemical inertness. Depending on deposition conditions, the properties of DLC films can be adjusted for several applications, e.g. to enhance the wear and corrosion resistance of precision cutting and machining tools. DLC is also used as a protective coating on magnetic hard disks and optical glasses (Laube et al. 2006). In a recent study (Laube et al. 2007), the role of DLC in reducing ureteral stent encrustation was investigated. The authors considered ten patients with heavy encrustations, different underlying diseases and stent removal frequency under 6 weeks due to encrustation and concluded that DLC coating improves the surface properties of ureteral stents and decreases encrustation tendencies and biofilm formation. The aim of our work was to evaluate the role of a heparin layer and a DLC layer applied to a ureteral stent in reducing double-J encrustations and we focused our attention more on stent surface analyses rather than on clinical aspects of ureteral stenting. For this purpose, FESEM and EDS were the most powerful techniques for characterizing stent surfaces. MicroIR spectroscopy analysis completed and confirmed data obtained from FESEM and EDS analyses. Using the micro-IR technique, only the nature

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of organic and inorganic groups of encrustations and of the polyurethane substrate can be recognized, making qualitative but not quantitative evaluations. Using the micro-IR technique, it is not possible to estimate the thickness of the biofilm, but this evaluation can easily be done with the electron microscope. In general, both organic and inorganic encrustations, derived from bacterial adhesion and urine flow, were formed on the coated stent after 1 month of insertion. Different kinds of deposit were detected on the surfaces of the stents: (a)

amorphous and crystalline inorganic deposits, like calcium oxalate, sodium chloride and brushite; (b) bacteria biofilm. The atomic compositions of these deposits were measured in percentage terms. In particular, by using the micro-IR technique, calcium oxalate could clearly be recognized because it presents a characteristic IR spectrum. It was possible to measure the thickness of the encrustations at the stent transversal cross-sections. It was observed that the biofilm layer was thicker on the inner surface of the stent than on the external one. This fact could be explained by assuming that a larger quantity of urine passes inside the lumen of the stent than outside it. In general, neither heparin-coated nor carbon-treated stents degraded over time; nor did they become stiffer and more brittle after 1 month of insertion. In the case of stents inserted for longer times, i.e. 10 and 12 months, the same results were found, confirming that neither polyurethane nor the surface treatments are chemically altered by contact with urine. The comparison of these data with those obtained from untreated polyurethane stents shows, in general, that the biofilm and inorganic encrustations deposited on the heparin- and carbon-coated catheters are thinner, and less compact and uniform than the encrustations formed on untreated stents. In particular, the presence of the surface treatment decreases the incidence of these encrustations, both outside and inside the lumen of the stent. In some cases, after the bilateral positioning of the surface-coated and uncoated stents, both of them showed remarkable levels of biofilm and inorganic encrustations, mostly caused by residual calculosis. What became clear from our clinical experience is that the use of a coated stent is the better choice in problematic patients, because it partially remains freer of encrustation compared with an untreated stent, which, in contrast, becomes completely covered with biofilm and inorganic deposits. The surface treatment enabled improved urine drainage and freeing of the stent lumen; nevertheless, a certain level of encrustation was still present. As already mentioned, clinical aspects of ureteral stenting were not a primary objective of this work; nevertheless, some consideration of these aspects could be useful. We did not record any technical problems during

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stent positioning or removal, showing no difference between a coated stent and traditional stent, and in all cases the procedure was simple and uneventful. During the study period, no patient reported fever or flank pain, and urine culture was negative in all cases; all these points being important for the use of these devices in a clinical setting. Overall, stents were well tolerated. No patients reported pain and the visual analog scale (VAS) score was very low in all cases. One patient with a heparin-coated stent (1/7, 14%) and two patients with carbon-coated stents (2/9, 22%) reported frequency/urgency symptoms. Obviously this difference was not significant for statistical analysis because of the small number of patients, so we cannot conclude which is the best stent as far as tolerance is concerned and further studies are required. The patients were successfully treated with anti-muscarinics until stent removal.

6.5

Conclusions

We have reported on some studies concerning the prevention of bacteria biofilm formation on polyurethane ureteral stents using two different surface treatments: heparin coating and amorphous carbon modification. Sixteen patients were studied. Five of them had heparin-coated stents and uncoated stents positioned bilaterally in their ureters, and another five patients had a carbon-coated stent and an untreated stent, for a period of 1 month. Two patients needed stenting for a prolonged period, so heparincoated stents were analyzed after 10 and 12 months of indwelling respectively. Finally, four patients, due to their pathology, received only the carbon-coated stent in one ureter for 1 month. After the indwelling time, the stents were removed and then characterized by FESEM, elemental analysis using EDS, and micro-IR spectroscopy. In conclusion, in all patients studied, the presence of the heparin layer or the carbon coating on the stent surface decreased the incidence of encrustation and the level of bacteria biofilm. When a compact layer of bacteria biofilm covered the surface of an untreated stent, the treated one, positioned in the same patient, showed a much lower level of biofilm and inorganic encrustation. The construction of a compact layer of salts, oxides and bacteria and its thickness were therefore reduced by the presence of a heparin or carbon layer, although the formation of biofilm and the precipitation of inorganic compounds, such as calcium oxalate and brushite, were not completely inhibited.

6.6

Acknowledgement

The financial support and stents supplied by Cook Ireland Ltd (Limerick, Ireland) are acknowledged.

Coated ureteral stents

6.7

155

References

appelgren p., ransjö u., bindslev l., espersen f., larm o. (1996). ‘Surface heparinization of central venous catheters reduces microbial colonization in vitro and in vivo: Results from a prospective, randomized trial’. Crit. Care Med. 24: 1482–1489. arshad m., shah s. s., abbasi m. h. (2006). ‘Applications and complications of polyurethane stenting in urology’. J. Ayub. Med. Coll. Abbottabad 18(2): 69–72. cauda f., cauda v., fiori c., onida b., garrone e. (2008). ‘Heparin coating on ureteral double J stents prevents encrustations: an in vivo case study’. J. Endourol. 22: 465–472. chew b. h., cadieux p. a., reid g., densted j. d. (2006). ‘In vitro activity of triclosan eluting ureteral stent against common bacterial uropathogens’. J. Endourol. 20(11): 949–958. cho y. h., lee s. j., lee j. y., kim s. w., kwon i. c., chung s. y., yoon m. s. (2001). ‘Prophylactic efficacy of a new gentamicin-releasing urethral catheter in short-term catheterized rabbits’. BJU Int. 87: 104–109. choong s. k., wood s., whitfield h. n. (2000). ‘A model to quantify encrustation on ureteric stents, urethral catheters and polymers intended for urological use’. BJU Int. 86: 414–421. cormio l., vuopio-varkila j., siitonen a., talja m., ruutu m. (1996). ‘Bacterial adhesion and biofilm formation on various double-J stents in vivo and in vitro’. Scand. J. Urol. Nephrol. 30: 19–24. cormio l., la forgia p., la forgia d., siitonen a., ruutu m. (2001). ‘Bacterial adhesion to urethral catheters: Role of coating materials and immersion in antibiotic solution’. Eur. Urol. 40: 354–358. costerton, j. w. (2007). The Biofilm Primer (Springer Series on Biofilms), Springer, Berlin. ditizio v., ferguson g. w., mittelman m. w., khourg a. e., bruce a. w., dicosmo f. (1998). ‘A liposomal hydrogel for the prevention of bacterial adhesion to catheters’. Biomaterials 19: 1877–1884. gorman s. p. and woolfson a. d. (2002). ‘Novel biomimetic and bioactive silicones’. Med. Device Technol. 13: 14–15. hildebrandt p., rzany a., bolz a., schaldach m. (1999). ‘Immobilisiertes heparin als inkrustierungsresistente. Beschichtung auf urologischen implantaten’. Biomed. Technol. 42: 123–124. imram m. a. (1999). Stent with silver coating and method. United States Patent 5976169. jansen b., goodman l. p., ruiten d. (1993). ‘Bacterial adherence to hydrophilic polymer-coated polyurethane stents’. Gastrointest. Endosc. 39: 670–673. john t., rajpurkar a., smith g., fairfax m., triest j. (2007). ‘Antibiotic pretreatment of hydrogel ureteral stent’. J. Endourol. 21(10): 1211–1216. kohler-ockmore j., feneley r. c. l. (1996). ‘Long-term catheterisation of the bladder: prevalence and morbidity’. Br. J. Urol. 77: 347–351. laube n., bradenahl j., meißner a., rappard j. v., kleinen l., müller s. (2006). ‘Plasma-deposited carbon coating on urological indwelling catheters: Preventing formation of encrustations and consecutive complications’. Urologe A. 45(9): 1163–1164.

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laube n., kleinen l., bradenahl j., meissner a. (2007). ‘Diamond-like carbon coatings on ureteral stents – a new strategy for decreasing the formation of crystalline bacterial biofilms?’ J. Urol. 177(5): 1923–1927. leung w. c., lau g. t., sung j. j., costerton j. w. (1992). ‘Decreased bacterial adherence to silver-coated stent material: an in vitro study’. Gastrointest. Endosc. 38: 338–340. leung j. w., liu y., cheung s., chan r.c., inciardi j. f., cheng a. f. (2001). ‘Effect of antibiotic-loaded hydrophilic stent in the prevention of bacterial adherence: A study of the charge, discharge, and recharge concept using ciprofloxacin’. Gastrointest. Endosc. 53: 431–437. maki d. g., cobb l., garman j. k., shapiro j. m., ringer m., helgerson r. b. (1988). ‘An attachable silver impregnated cuff for prevention of infection with central venous catheter: a prospective randomized multicenter trial’. J. Am. Med. Assoc. 127: 267–274. raad i., darouiche r., dupuis j., abi-said d., gabrielli a., hachem r. (1997). ‘Central venous catheters coated with minocycline and rifampin for the prevention of catheter-related colonization and bloodstream infections’. Ann. Intern. Med. 127(4): 267–274. reid g., habash m., vachon d., denstedt j., riddell j., beheshti m. (2001). ‘Oral fluoroquinolone therapy results in drug adsorption on ureteral stents and prevention of biofilm formation’. Int. J. Antimicrob. Agents 17: 317–319. riedl c., witkowski m., plas e., pflueger h. (2002). ‘Heparin coating reduces encrustation of ureteral stents: a preliminary report’. Int. J. Antimicrob. Agents 19: 507–510. riley d. k., classen d. c., stevens l. e., burke j. p. (1995). ‘A large randomized trial of a silver impregnated urinary catheter: lack of efficacy and staphylococcal superinfection’. Am. J. Med. 98: 349–358. ruggieri m. r., hanno p. m., levin r. m. (1987). ‘Reduction of bacterial adherence to catheter surface with heparin’ J. Urol. 138: 423–426. sabbuba n., hughes g., stickler d. j. (2002). ‘The migration of Proteus mirabilis and other urinary tract pathogens over Foley catheters’. BJU Int. 89: 55–60. saint s., elmore j. g., sullivan s. d., emerson s. s., koepsell t. d. (1998). ‘The efficacy of silver alloy-coated urinary catheters in preventing urinary tract infection: a meta-analysis’. Am. J. Med. 105: 236–241. stickler d. j. (2000). ‘Biomaterials to prevent nosocomial infections: is silver the gold standard?’ Curr. Opin. Infect. Dis. 13: 389–393. stickler d. j., morris n. s., williams t. j. (1996). ‘An assessment of the ability of a silver-releasing device to prevent bacterial contamination of urethral catheter drainage systems’. Br. J. Urol. 78: 579–588. tenke p., reidl c. r., jones g. l., williams g. j., stickler d., nagy e. (2004). ‘Bacterial biofilm formation on urologic devices and heparin coating as preventive strategy’. Int. J. Antimicrob. Agents 23S: 67–74. warren j. w., muncie h. l., hebel j. r., mall-craggs m. (1994). ‘Long-term urethral catheterization increases risk of chronic pyelonephritis and renal inflammation’. J. Am. Geriatr. Soc. 42: 1286–1290. watterson j. d., cadieux p., denstedt j. d. (2002). ‘Ureteral stents: which, when, and why?’ AUA Update Ser. 21: 122–127.

7 Proteus mirabilis biofilm formation and catheter design D. J. S T I C K L E R, Cardiff University, UK

Abstract: This chapter describes the particular problems that infections with Proteus mirabilis cause patients undergoing long-term indwelling bladder catheterization. The mechanisms by which this organism colonizes catheter surfaces and forms the crystalline bacterial biofilms that encrust and block catheters are discussed. The reasons for the vulnerability of the currently available devices to these problems are explained. Suggestions for novel devices are reviewed and it is concluded that there is plenty of scope to improve both the manufacturing standards and the design of these catheters. Key words: Proteus mirabilis, catheter encrustation, urolithiasis, Foley catheters, bacterial biofilms.

7.1

Introduction

Proteus mirabilis is an extraordinary microbe. It was named Proteus after an elusive character in Homer’s Odyssey who was able to escape capture by changing its form. There are other species in the genus, e.g. the common Proteus vulgaris, but the one of clinical relevance is called mirabilis, from the Latin meaning wonderful or surprising. The very name thus tells us that it is a special bacillus. Its natural habitat is the intestinal tract. It is part of the faecal flora of humans and many other mammals. It can be isolated from sewage or polluted soils and waters with ease. It is not a primary invasive pathogen and rarely causes infection in normal healthy individuals. It is an opportunist, given a debilitated host it can produce infection. It is rarely responsible for urinary tract infections in normal, otherwise healthy individuals. In patients with urinary tracts subject to chronic instrumentation, however, it becomes a significant pathogen. Unfortunately for us it has many properties that make it ideally suited to life in the catheterized urinary tract (Mobley, 1996). 157

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7.2

Virulence factors

P. mirabilis has an arsenal of virulence factors that facilitate colonization and subsequent infection of a debilitated host. It is a very sticky bacillus having at least four different adhesins that mediate its attachment to tissue surfaces (Rocha et al., 2007). It has an exopolysaccharide capsule that protects it against hosts’ defences and helps it to bind firmly to surfaces. It secretes a haemolysin, an iron scavenging protein, proteases and amino acid deaminases – all of which are essential for extracting important nutrients from host tissues and fluids. A specific immunoglobulin A (IgA) protease that it produces is capable of degrading the predominant immunoglobulin in the mucus secreted from epithelial surfaces (Jacobsen et al., 2008). In addition it has two signature features that are particularly important in the catheterized urinary tract: the ability to produce the enzyme urease and the ability to swarm rapidly over surfaces (Belas, 1996; Mobley, 1996).

7.2.1 Urease The urease of P. mirabilis is a particularly potent enzyme, being able to hydrolyse urea several times faster than the ureases produced by other species (Jones and Mobley, 1987). It is a cytoplasmic nickel-metalloprotein induced by its substrate urea. As urea is always present in human urine in concentrations up to 500 mM, the enzyme is in fact constitutively expressed in the presence of urine. Urea passes freely across the bacterial cell membrane and is rapidly hydrolysed to yield two molecules of ammonia and one of carbon dioxide. Reaction with water results in the formation of two ammonium and hydroxyl ions to one each of bicarbonate and hydrogen. The net result is that the urine becomes more alkaline. As the pH of the urine rises, calcium and magnesium phosphates come out of solution and form crystals of struvite (magnesium ammonium phosphate) and carbonateapatite (calcium carbonate phosphate). Urolithiasis develops as aggregations of the crystalline material and bacterial cells generate bladder and kidney stones and encrust any devices implanted in the infected urinary tract (Griffith et al., 1976; Dumanski et al., 1994; Morris et al., 1999). The role of urease in the pathogenicity of P. mirabilis infections has also been demonstrated in experimental animal models. A urease-negative mutant was found to be less able to colonize the bladder and kidneys of mice than the wild-type parent strain. The urease-positive parent strain also produced significantly more severe renal pathology than the mutant (Johnson et al., 1993a).

7.2.2 Swarming Microscopic examination of samples of urine infected with P. mirabilis will reveal the presence of small Gram-negative bacilli about 2 μm in length.

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They swim around quite actively being propelled by the action of a small number (1–10) of flagella (Fig. 7.1(a) and (b)). When these cells attach to a surface, however, spectacular changes can occur which transform them into cells that can be up to 80 μm long. These elongated cells can produce up to 5000 flagella per cell. They no longer appear as single cells, they organize themselves into parallel groups, tied together by the helical binding of the flagella around adjacent cells (Fig. 7.1(c) to (e)). They are then capable of moving rapidly in a co-ordinated fashion (swarming) over solid surfaces (Williams and Schwarzhoff, 1987; Mobley and Belas, 1995; Jones et al., 2004). In this way they can extend their populations over surfaces and colonize new locations. These swarmer cells have also been shown to produce elevated levels of urease, haemolysin and proteases (Falkinham and Hoffman, 1984; Allison et al., 1992). The phenomenon of swarming can be observed on ordinary nutrient agar plates. On incubation of a suspension of the normal vegetative bacilli on agar, the single cells multiply until a high population density is reached. Fundamental changes then occur in their genetic expression involving the co-ordinated switching of many genes (Rather, 2005). As a result, cells at the edge of the micro-colony elongate, become hyperflagellated and organize themselves into highly motile multi-cellular rafts that move out from the original inoculation site. They also cover themselves in an extracellular slime that reduces surface frictional forces and facilitates migration (Stahl et al., 1983). This swarming process stops after a while and the cells revert to the small vegetative cells with few flagella that can no longer move over the surface. After a period termed ‘consolidation’, in which the population builds again, swarmer cells are regenerated and migrate out once more. This process is replicated until the surface of the agar plate is totally covered with characteristic rings of growth (Fig. 7.2(a) and (b)).

7.2.3 Migration over catheter surfaces Given its ability to swarm rapidly and spread over solid surfaces and the fact that urease activity in swarmer cells is induced to levels at least 30-fold that in normal bacilli (Allison et al., 1992), it is not surprising that it has been suggested that swarming may have important roles in the pathogenesis of P. mirabilis catheter-associated infections (Belas, 1996). Stickler and Hughes (1999) showed that P. mirabilis could swarm rapidly over allsilicone, silicone-coated latex, hydrogel-coated latex and hydrogel/silvercoated latex catheters. Migration was significantly more rapid over the two hydrogel-coated catheters. Electron microscopy revealed the presence of discrete rafts of elongated swarmer cells on catheter surfaces (Fig. 7.2(c) and (d)). It was suggested that swarming plays a part in: (a) the initiation of the infections, facilitating the migration of P. mirabilis from the insertion

(b)

2 μm

Det WD Exp SE 13.6 17 2B431

Acc.V Spot Magn Det WD Exp 30.0 kV 2.0 11210x SE 11.1 55 B4 RHT

(e)

1 μm

(c)

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7.1 Electron micrographs of P. mirabilis showing the normal cells swimming in an aqueous medium (a) and (b). The short hair-like structures visible in panel (a) are fimbrial adhesins. The cells are motile owing to the action of a small number of flagella that are visible in panel (b). The leading edge of a swarming front on agar is shown in panel (c). It can be seen that the organism has transformed itself into long multi-flagellated cells arranged in parallel masses. In panels (d) and (e) the swarmer cells are shown to be tied together by the helical binding of flagella around adjacent cells.

Acc.V Spot Magn Det WD Exp 30.0 kV 2.0 31130x SE 13.6 17 2B433

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Proteus mirabilis biofilm formation and catheter design (a)

(b)

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Spot Magn Det WD Exp 50 μm 3.0 800x SE 10.7 1 o9sD Hydreget couted latex

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7.2 Image (a) shows a Petri dish containing two blocks of nutrient agar that have been bridged by three 1 cm sections of catheter. Drops of a suspension of P. mirabilis were inoculated adjacent to one end of each of the catheter sections. The plate was incubated at 37 °C for 18 h and the result is shown in panel (b) (from Stickler and Hughes (1999), with permission). It can be seen that the bacteria have swarmed on the agar, over the catheter bridges and continued swarming on the agar on the other side of the bridge. Panels (c) and (d) are scanning electron micrographs showing rafts of swarmer cells moving from left to right over the irregular surface of a hydrogelcoated latex catheter section that had been removed from the plates before the test organism had reached the other side (from Stickler (2005), with permission).

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site along the catheter into the bladder, and (b) the spread of the crystalline biofilm over the catheter. A subsequent study showed that P. mirabilis was able to migrate over all the basic types of catheter more effectively that any other urinary tract pathogen (Sabbuba et al., 2002). It was significant that when these tests were performed under conditions in which swarming did not occur, migration over catheters was impeded. Mutants lacking the ability to swarm also failed to migrate over catheters (Jones et al., 2004). The facility with which P. mirabilis moves over catheters compared with other pathogens, suggests that swarming has a role in the initiation of catheter-associated urinary tract infection by facilitating the movement of P. mirabilis from the skin at the catheter insertion site, along the outside of the catheter into the bladder. It might also explain why an organism present in faecal flora is a much more important pathogen of the catheterized tract than of non-catheterized individuals. Recently Goanakar et al. (2007) demonstrated that catheters impregnated with chlorhexidine and triclosan inhibit the migration of P. mirabilis. It would be interesting to test whether these catheters could prevent infection with P. mirabilis in patients.

7.3

Epidemiology of Proteus mirabilis infections

P. mirabilis is not generally among the pioneer colonizers of the catheterized urinary tract and is not commonly found infecting patients undergoing short-term catheterization (Matsukawa et al., 2005). The longer the catheter is in place, the more likely P. mirabilis is to colonize the urine. In patients undergoing long-term catheterization it has been isolated from 44% of urine samples (Warren et al., 1982). While a number of other ureaseproducing species (e.g. Morganella morganii, Klebsiella pneumoniae and Pseudomonas aeruginosa) can be responsible for catheter-associated urinary tract infection, there is strong epidemiological and experimental evidence that P. mirabilis is the main cause of the complication of catheter encrustation. It is the organism most commonly isolated from the urine of patients suffering from recurrent catheter encrustation and blockage (Mobley and Warren, 1987; Kunin, 1989). It was also the species most commonly recovered from patient’s encrusted catheters (Stickler et al., 1993). Experimental work in a laboratory model demonstrated that M. morganii, K. pneumoniae and Ps. aeruginosa failed to produce alkaline urine and generate appreciable encrustation on catheters (Stickler et al., 1998). The only species capable of producing alkaline urine and causing extensive encrustation were P. mirabilis, P. vulgaris and Providencia rettgeri. These latter two species are not commonly found in the urine of catheterized patients, so clearly P. mirabilis is the main cause of the problem.

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Sabbuba et al. (2003) developed a method of genotyping P. mirabilis to facilitate a better understanding of the epidemiology and pathogenesis of catheter-associated urinary tract infections with this species. Pulsed-field gel electrophoresis of restriction enzyme digests of P. mirabilis DNA produced highly discriminatory genotypic profiles. Application of this method established the remarkable stability of and persistence of strains of P. mirabilis in the catheterized urinary tract; the same genotype persisting in a patient’s urinary tract despite many catheter changes, courses of antibiotic treatment and even periods when the patient was not catheterized. It was suggested that this stability is due to the presence of P. mirabilis in the bladder stones that frequently form in these patients. Subsequent genotyping of pairs of P. mirabilis isolates from the encrusted catheters and bladder stones of patients demonstrated that in each case the isolate from the stone was indeed identical to that from the catheter (Sabubba et al., 2004). Although P. mirabilis is generally sensitive to most antibiotics, it is difficult to clear from the catheterized urinary tract by antibiotic treatment. The probable reason for this is that bacteria within the matrix of the bladder stones are protected from antibacterial agents. It is thus clear that the bladder stones harbour the strains of P. mirabilis and ensure the rapid colonization of replacement catheters by crystalline biofilm. During this work it also became clear that there was little evidence of cross-infection with P. mirabilis especially between the community-based patients. All these patients were infected with genetically distinct strains. Subsequent analysis showed that pairs of faecal and catheter biofilm isolates from patients were identical (Mathur et al., 2005). The hypothesis is thus that most long-term catheterized patients who suffer from catheter encrustation acquire P. mirabilis from their own faecal flora. These strains will eventually chronically colonize the bladder and ensure their persistence in the urinary tract in the form of bladder stones.

7.4

The process of crystalline biofilm formation on catheters

It is important to appreciate precisely how P. mirabilis colonizes, encrusts and blocks catheters, if we are to devise effective strategies for the control of the problem. In simple terms the bacteria adhere to the catheter, then bathed in a gentle flow of warm nutritious urine, they multiply rapidly producing enormous populations of cells and form extensive bacterial biofilms embedded in their protective gel-like matrix of polysaccharide. While this process is going on, the urease activity generates alkaline urine and this in turn causes the precipitation of crystals of calcium and magnesium phosphates in the urine and the biofilm. The resulting crystalline biofilm

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eventually blocks the flow of urine through the catheter (Morris et al., 1999). In recent years it has become obvious that P. mirabilis is an ingenious organism capable of initiating the development of these crystalline biofilms in a variety of ways. It is thus able to colonize a variety of different types of catheter under a range of environmental conditions. The general model for the development of biofilms on implanted prosthetic devices involves the rapid coating of the device with a conditioning film of host proteins from the surrounding body fluids. These proteins are considered to provide receptors for the adhesins present on the bacterial surface. Adherence to the surface is secured through binding of the adhesins to the receptors. The cells then secrete exopolysaccharide which fixes the cells to the surface. Subsequent bacterial multiplication within the polysaccharide gel generates the biofilm (Donlan and Costerton, 2002). This probably happens with P. mirabilis on urinary catheters. The cells have a number of different adhesins, fine hair-like fimbrae that protrude from their surface (Rocha et al., 2007). There is evidence that these structures are involved in the attachment of the bacteria to epithelia and the initiation of infection in the urinary tract (Bahrani et al., 1994). Examination of catheters removed from patients after short periods of time have also revealed coatings of proteins such as fibrin (Ohkawa et al., 1990). Santin et al. (1997) also reported that proteins including albumin and α-1-microglobulin were present on the surfaces of ureteral stents removed from patients. There is evidence from experimental work, however, that P. mirabilis cells can also bind directly on to surfaces such as silicone from suspensions in buffered saline (Downer et al., 2003a). It seems therefore that the bacilli can bind to catheters whether or not they are coated in host proteins. It is important to recognize that in addition to the biological adherence factors, there are also powerful physical and chemical forces involved in the initiation and development of the crystalline biofilms that block catheters. Scanning electron microscopy on unused catheters has revealed the rough, irregular nature of their surfaces (Axelsson et al., 1977; Cox, 1990). Latexbased catheters have particularly uneven surfaces (Stickler et al., 2003a) especially around the eye-holes where the manufacturing techniques tear through the latex and produce surfaces on the rims of the eye-holes that must look like a rocky landscape of craters and crevices to a bacterium (Fig. 7.3). The common occurrence of diatom skeletons embedded in the lumenal surfaces, exacerbates the surface roughness of latex catheters. The source of these silica skeletons is the diatomaceous earth that is used to prevent the catheter material sticking to the metallic formers on which catheters are produced. All-silicone catheters have rather smoother surfaces but irregularities are still common where the eye-hole has been gouged out of the silicone tubing and where extrusion manufacturing techniques have produced striations on the lumenal surfaces (Fig. 7.4).

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7.3 Scanning electron micrographs showing the surfaces of unused latex-based catheters. Panels (a) and (b) are images of the eyelets of a hydrogel-coated latex catheter. The silica skeletons of diatoms can be seen on the irregular lumenal surfaces of a hydrogel-coated latex catheter (c) and a silver/hydrogel-coated latex catheter (d).

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7.4 Scanning electron micrographs showing the surfaces of unused all-silicone catheters. Panels (a) and (b) are images of the catheter eyelets; (c) and (d) show the striated nature of the lumenal catheter surface.

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As contaminated urine flows over the catheter surfaces, cells can be trapped in the crevices, craters and other irregularities. Experiments in physical models of the catheterized bladder in which catheters were removed at various time intervals after infection with P. mirabilis, demonstrated the particular vulnerability of the eye-holes to colonization (Stickler et al., 2003a). Scanning electron microscopy revealed that within 2 h, bacterial cells were trapped in the crevices in the uneven surfaces of the eyelets. Micro-colonies of cells developed in the surface depressions and then, with the rise in urinary pH, crystals started to form in the biofilm. Extensive crystalline biofilm developed and spread down the catheter lumen (Fig. 7.5). It was also obvious that the silica skeletons of the diatoms embedded in the latex were attractive sites for bacterial colonization. Blockage with extensive crystalline biofilm generally occurred at the eye-hole or in the balloon region of the lumen (Fig. 7.6). Experiments in parallel-plate flow cells showed that when urine cultures flow over smooth flat polymer films, the pH of the urine can be a major factor in determining bacterial adhesion. For example, some polymers with strongly electron donating surfaces will resist colonization by cells until the pH of the urine rises above the pH at which calcium and magnesium phosphates crystallize out of solution. In the alkaline urine, macroscopic aggregates of cells and crystals form, settle on the polymer surface and initiate crystalline biofilm formation (Stickler et al., 2006a). In patients who suffer from the complication of catheter encrustation the usual management policy is just to change the blocked catheters. Under these circumstances fresh catheters are placed directly into urine cultures of P. mirabilis at alkaline pHs, containing micro-crystals of the calcium and magnesium phosphates. Laboratory investigations designed to examine the early stages of crystalline biofilm formation under these conditions revealed a common sequence in the development of the crystalline biofilm on allsilicone, silicone-coated latex, hydrogel-coated latex and hydrogel/silvercoated latex catheters (Morgan, 2007). After only 1 h in the models the catheter surfaces were covered by a micro-crystalline layer. X-ray microanalysis confirmed that this material was composed largely of calcium and phosphate. Bacterial colonization of this foundation layer followed with micro-colonies of cells forming on the crystals. The microcrystalline foundation layer has also been observed on catheters removed from patients (Fig. 7.7). By 18 h the eyelets and lumenal surfaces of all these catheters were comprehensively covered by densely populated crystalline P. mirabilis biofilm.

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7.5 Scanning electron micrographs illustrating how biofilms can be formed by P. mirabilis cells being trapped in surface irregularities of hydrogel-coated latex catheters. Panel (a) shows cells caught in a depression in the rim of an eye-hole. Cells forming a micro-colony on a diatom skeleton that was embedded in the lumenal surface are shown in panel (b) (from Stickler and Morgan (2008), with permission). The spreading of micro-colonies over the eyelet and the subsequent formation of the crystalline biofilm can be seen in panels (c) and (d).

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7.6 Low-vacuum scanning electron micrographs showing examples of crystalline P. mirabilis biofilms occluding the eye-holes and central channels of silver/hydrogel-coated latex catheters ((a) and (c)) and nitrofurazone-coated silicone catheters ((b) and (d)) (from Stickler and Sabbuba (2008), with permission).

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7.7 Scanning electron micrographs showing the early stages of biofilm formation on catheters placed in bladder models containing cultures of P. mirabilis in urine at pH 8.5 ((a) and (b)). The first stage in biofilm development under these conditions is the formation of a foundation layer composed of micro-crystals of calcium phosphate on to which bacteria adhere (a) and then biofilm develops (b). Panel (c) shows the micro-crystalline structure of the foundation layer on which biofilm is developing on a silver/hydrogel-coated latex catheter removed from a patient.

7.5

Antimicrobials in the prevention of catheter encrustation

The observations that crystalline biofilms can develop on the foundation layer of calcium phosphate and by the deposition from urine of macroscopic co-aggregates of cells and crystals, have profound implications for the development of encrustation-resistant catheters. Catheters with surface properties aimed at inhibiting bacterial attachment and biofilm development by immobilizing an antibacterial in the surface are thus unlikely to be effective in preventing encrustation in patients infected with P. mirabilis. In

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the case of the silver catheters for example, the foundation layer allows cells to attach and grow protected from contact with the underlying silver. The clear lesson is that in order to prevent catheter encrustation it is essential to prevent the pH of the urine rising above the pH at which crystals form. If antimicrobials are to be incorporated into catheters to achieve this, they must diffuse out from the catheter surface and reduce the viable cell populations of P. mirabilis in the urine. In the case of the silver/hydrogel-coated latex catheter, insufficient silver elutes into the urine to inhibit the activity of P. mirabilis. An alternative silver catheter produced from silicone impregnated with nano-particulate silver has recently been described (Samuel and Guggenbichler, 2004). It will be interesting to know whether this catheter is capable of producing urinary concentrations of silver that are bactericidal. Chakravarti et al. (2005) showed that it was possible to generate urinary concentrations of silver that inhibited P. mirabilis-induced encrustation by passing an electric current through silver electrodes attached to catheters. The effect was temporary, however, as the silver electrodes disintegrated after about 150 h. Observations on the colonization of nitrofurazone-coated silicone catheters in bladder models (Morgan, 2007) suggested that little encrustation takes place in the first 4 h. At 12 h, however, cells could be seen attaching both directly to the catheter and to areas where micro-crystalline aggregates had formed. At 18 h the catheter had acquired a mature crystalline biofilm comparable with those formed on the all-silicone catheters. By 29 h the eye-holes and central channels were blocked by extensive crystalline deposits (Fig. 7.6). In these tests with the nitrofurazone catheters, the pH of the urine in these models at 4 h had fallen to 6.4, below the level at which phosphates crystallized from this urine. By 18 h, however, the pH had recovered to 8.3. The viable cell count of the urine also fell until 18 h. During these experiments it was observed that the yellow colour characteristic of nitrofurazone, eluted rapidly from these catheters over the first few hours. Nitrofurazone is not very active against P. mirabilis, concentrations of 32–64 μg/ml are required to inhibit its growth (Johnson et al., 1993b). It is possible that high concentrations of the drug eluted initially into the urine and had an antibacterial effect but the cultures recovered, alkaline conditions were re-established and catheter encrustation was resumed as the drug concentration decreased with time. Given the problem of maintaining the release of effective concentrations of antibacterials into urine for the life time of long-term catheters, Bibby et al. (1995) put forward the suggestion that the retention balloon could be used as a reservoir for antibacterial agents. The idea being that the membrane of the balloon might ensure the controlled release of the active agent over long periods. These workers were able to show that mandelic acid could diffuse through the catheter balloon producing concentrations of around 0.1 mg/ml in the surrounding urine. Unfortunately mandelic acid

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has poor activity against P. mirabilis and other urinary tract pathogens, being bactericidal in urine only at concentrations of above 5 mg/ml (Rosenheim, 1935). In contrast to mandelic acid, the biocide triclosan is particularly active against P. mirabilis. The minimum inhibitory concentration (MIC) of this agent against strains of P. mirabilis isolated from encrusted catheters has been reported to be less than 0.5 μg/ml (Stickler, 2002). Using laboratory models supplied with artificial urine and infected with P. mirabilis, Stickler et al. (2003b) demonstrated that triclosan can diffuse through the balloons of all-silicone catheters, significantly reduce the number of viable cells in the urine, prevent the rise in urinary pH and inhibit crystalline biofilm formation on the catheters. In these experiments control catheters with their retention balloons inflated with water, blocked within 24 h, whereas the test catheters inflated with triclosan (10 mg/ml in 5% w/v polyethylene glycol) drained freely for the 7 day experimental period. Examination of these catheters revealed little sign of encrustation at the eye-hole or in the lumen. It is thus likely that they would have drained freely for considerably longer than 7 days. Subsequent experiments showed that this strategy was also effective with latex-based catheters and when artificial urine was replaced with pooled human urine. Triclosan was found to have impregnated the whole length of the all-silicone catheters, but not the latex-based devices (Jones et al., 2005). The conditions used in these experiments were intended to simulate those in which a catheter is introduced into residual bladder urine that is heavily colonized by P. mirabilis (108 colony forming units (cfu)/ml) at a pH of around 8.5. A concentrated urine and a low flow rate through the bladder were also employed to simulate the low fluid intake characteristic of many elderly patients undergoing long-term catheterization. The observation that encrustation is inhibited under these severe experimental conditions that normally produce rapid catheter blockage, suggests that the triclosan strategy should extend substantially the life-span of catheters in patients prone to encrustation problems. Chemical analysis demonstrated that loading the retention balloons with triclosan (10 mg/ml in 5% w/v polyethylene glycol) results in the daily diffusion of around 115 μg of the agent into the urine (Jones et al., 2006). Assuming this diffusion rate is maintained, the antibacterial activity should persist in the urine for well over the current maximum life-span of 12 weeks for each long-term catheter. The strategy was also found to inhibit biofilm formation by Escherichia coli, Klebsiella pneumoniae and Providencia stuartii. Not all of the species that can cause catheter-associated urinary tract infection are sensitive to triclosan however. Pseudomonas aeruginosa, Morganella morganii and Serratia marscens (MICs of triclosan >100 mg/l) were able to colonize the triclosan-impregnated silicone catheters (Jones et al., 2006). These species however, despite being urease positive in laboratory

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bacterial identification tests, do not seem to be able to produce alkaline urine or crystalline biofilms (Mobley and Warren, 1987; Stickler et al., 1998). If the triclosan strategy could be successfully transferred from the laboratory to the clinic, it could bring major improvements in the bladder management of many elderly and disabled people. The strategy of delivering agents directly to the bladder via the catheter balloon, does not disturb the integrity of the closed drainage system and does not require the manufacture of novel catheters. It could possibly be used to deliver other agents. Carlsson et al. (2005), for example, showed that nitric oxide can diffuse through catheter balloons by inflating the balloons with a mixture of sodium nitrite and ascorbic acid. Williams and Stickler (2007) screened the ability of 18 biocides and antibiotics to diffuse through silicone balloons. They reported that only triclosan and nalidixic acid were capable of diffusing through the balloons in sufficient concentrations to inhibit the encrustation process significantly. There is surely scope, however, to manipulate the composition of the polymers used in the manufacture of catheter balloons to modify their diffusion properties. Polyurethane balloons, for example, were shown to be far more permeable than silicone balloons to gentamicin and the fluoroquinolones (Williams and Stickler, 2007) and could be considered as an alternative to silicone or latex in catheter manufacture. As with any intervention involving an antibacterial agent, the possibility of resistance developing is a concern. In the case of triclosan, despite extensive use in many preparations for over 30 years, there has been little sign that the clinical or domestic use has led to the generation of resistant organisms (Sreenivasan and Gaffar, 2002; McBain et al., 2003; Russell, 2004). Recently, however, it has been reported that exposure to triclosan under laboratory conditions can result in the development of strains of P. mirabilis with reduced susceptibility to the biocide. A strain with an MIC of 40 mg/l was able to encrust and block catheters that had been primed with triclosan (Stickler and Jones, 2008). In clinical trials and any subsequent clinical application of this strategy, it will be particularly important to monitor the urinary flora for signs of resistance to triclosan. The risk of developing resistance could be reduced if triclosan was not used prophylactically, by replacing water for the inflation of all catheter balloons, but as an intervention when the urine becomes colonized by P. mirabilis and encrustation becomes a problem. The sensor developed to give an early warning of impending catheter encrustation and blockage by signalling that P. mirabilis has infected the urine, would permit the implementation of such a strategy (Stickler et al., 2006b). These sensors are simple plastic strips composed of cellulose acetate to which the acid/akali indicator bromothymol blue has been covalently conjugated. They are designed to be placed in the drainage system so that they are clearly visible to the patient and carer. They

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continuously monitor the pH at their surface. The concept is that a sustained change from the acid to alkaline reaction at the sensor surface will signal infection and biofilm formation by P. mirabilis or related urease-producing organisms. The appearance of the signal indicates that action should be taken to avoid an acute clinical episode. Experiments in a laboratory model of the catheterized bladder (Stickler et al., 2006b) and a study in catheterized patients (Stickler et al., 2006c), demonstrated that the sensors gave clear simple signals, changing in colour from yellow to dark blue in the presence of P. mirabilis. The indicator did not elute from the sensor and the urine drainage bag was found to be the optimum location for the sensor. Experiments in laboratory models showed that at this site the devices gave strong signals at least 40 h before the catheters blocked. Scanning electron microscopy confirmed that when the sensor in the drainage bag gave the signal, crystalline biofilm had started to form at the eye-hole region of the catheter and on the sensors themselves. The study in patients confirmed that the sensors were able to signal the early stages of catheter encrustation in good time to allow catheter replacement, thus avoiding the clinical crises induced by catheter blockage. A modulation strategy such as the use of triclosan in the balloon could thus be implemented in those cases where the sensor gives its signal. The idea is that the sensors could be placed in the drainage bags of all patients undergoing long-term catheterization from the time of their first catheter. Usually P. mirabilis is not among the early colonizers of the catheterized urinary tract, it tends to colonize later on and eventually can be found in about 40% of these patients’ urines (Mobley, 1996). When the sensor signals P. mirabilis infection, the strategy could be deployed. The continued use of the sensor could then monitor the efficacy of the strategy and signal any recurrences of the problem. Once P. mirabilis has established itself in the catheterized urinary tract and is producing urolithiasis, it is difficult to eliminate with antibiotic treatment (Subbuba et al., 2003; Hung et al., 2007). It is clearly not a harmless organism in the catheterized urinary tract. Left in peace, it will produce many complications in the care of these patients. Surely efforts should be made to prevent P. mirabilis establishing itself in the bladder and producing the stones that will secure its refractility to antibiotic therapy and ensure its permanent residence in the urinary tract. There is a strong argument that such an organism should be targeted with antibiotic therapy as soon as it appears in the catheterized tract. The sensor placed in the catheter drainage systems of patients from the time of their first catheter, could provide continuous monitoring of the urine for the presence of P. mirabilis. When the sensor turns blue, bacteriological analysis of the urine could then examine the antibiotic susceptibility of the invading strain and appropriate drugs prescribed before the organism protects itself in the crystalline aggregates that generate all the problems.

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Factors that modulate the rate of Proteus mirabilis biofilm formation on catheters

While the basic mechanism of catheter encrustation is well established, little is known about the factors that modulate the rate of crystalline biofilm development. Nurses caring for catheterized patients are aware that for patients designated as recurrent blockers, the time taken for the catheters to block (catheter life-span) varies considerably. The results of a prospective study on patients who were all infected with P. mirabilis, revealed that the time taken for catheters to block varied from 2 to 98 days (Mathur et al., 2006a). As encrustation is brought about by the rise in urinary pH caused by urease activity, it would seem likely that the higher the pH generated in the urine by P. mirabilis, the shorter would be the catheter life-span. The results of the study, however, suggested a more subtle relationship between the urinary pH and the encrustation process.

7.6.1 The nucleation pH of urine and the rate of encrustation If you take any sample of normal urine and gradually increase its pH, there will come a point, termed the ‘nucleation pH’ (pHn), at which it will become turbid due to micro-crystals of apatite and struvite coming out of the solution. Choong et al. (2001) showed that it is the relationship between the pH of the urine voided by catheterized patients and the pHn of that urine that determines whether the patients’ catheters will become encrusted. Choong et al. found that for patients designated as blockers, the mean pH of the voided urine was 7.85 while the mean pHn of their urine was 7.58, clearly explaining why crystals form in their urine. The data from the Mathur et al. (2006a, 2006b) studies showed that in the patients infected with P. mirabilis, the pHn of the urine, rather than its actual pH, is the important factor in predicting the rate of catheter encrustation. Statistical analysis revealed a highly significant (P = 0.004) correlation between pHn and catheter life-span. The higher the pHn value, the slower the rate of encrustation and the longer catheters took to block. The study also demonstrated that while there were individuals with consistently high or low pHn values, in some subjects the value varied considerably from week to week. These observations suggested that it should be possible to manipulate pHn as a therapeutic approach to deal with catheter encrustation. In a study with healthy volunteers (Suller et al., 2005), it was demonstrated that simply by increasing fluid intake and the citrate content of urine, its pHn could be elevated to values that are rarely achieved in P. mirabilis-infected urine. Subsequent experiments (Stickler and Morgan, 2006) in a laboratory model infected with P. mirabilis showed

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that when the bladder was supplied with a concentrated urine (pHn 6.7) at a low fluid output (720 ml/24 h), typical of many elderly catheterized patients, catheters blocked at a mean time of 24 h. Diluting this urine 1 in 4 increased the pHn to 7.5 and models supplied with this urine at 2880 ml/24 h took mean times of 125 h to block. When models were supplied with urine containing 0.4 mg/ml citrate at 1440 ml/24 h, catheters blocked at mean times of 66 h. When the citrate concentration of the urine was increased to 1.5 and 3.0 mg/ml (the pHn of urine rising to 8.3 and 9.1 respectively), the catheters drained freely for the full 7 day experimental period. Scanning electron microscopy revealed that the catheter biofilms that developed in urine with high pHn values were devoid of crystalline formations (Fig. 7.8). The advice to patients to increase their fluid intake by drinking steadily throughout the day (Burr and Nuseibeh, 1997) clearly has a sound basis in physiology and physical chemistry. The dilution of urine resulting from an increased fluid intake will elevate its pHn and slow the rate of catheter encrustation. If the citrate content of urine can also be elevated by encouraging patients to take lemon-based drinks for example, the rate of crystal formation should reduce further, even in P. mirabilis-infected urine. It is interesting that increasing a patient’s fluid intake with citrate-containing drinks (lemonade) has been advocated for the control of the related problem of infection-associated kidney stones (Seltzer et al., 1996). A recent paper (Kang et al., 2007) has reported no adverse side effects in patients who had been undergoing this therapy for a mean duration of 3 years. These observations should encourage a clinical trial to examine the effect of increasing a patient’s fluid intake with citrate-containing drinks on the encrustation and blockage of their catheters.

7.6.2 Biological control of encrustation In long-term patients, while infections with pure cultures of Proteus can be found, the urinary flora is usually a mixed population. Up to seven species have been recovered from some patients’ urine (Clayton et al., 1982). Little is known about the interaction of the different species in the catheter biofilms. The presence of other species in the catheterized urinary tract might affect the rate at which catheters encrust. Over the years, a database of the composition of biofilms recovered from patients’ catheters has been built up by successive workers in the catheter research laboratory of the Cardiff School of Biosciences, UK. Single organisms were found on 30 of the catheters, the remaining 76 were colonized by 2 or more species. Ps. aeruginosa, Enterococcus faecalis, E. coli and P. mirabilis were the most commonly found organisms. Similar results have been reported by other groups (Ohkawa et al., 1990).

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(a)

Acc.V Spot Magn Det WD Exp 5 μm 20.0 kV 4.0 10000x SE 13.1 5 A-0.41 gL citrate 24h (b)

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7.8 Scanning electron micrographs of P. mirabilis biofilms on catheters removed from laboratory models of the bladder after 24 h incubation. Panel (a) is a biofilm formed in urine containing 0.4 mg/ml of citrate with a nucleation pH of 7.4. Extensive micro-crystalline material typical of calcium phosphate can be seen throughout the biofilm. Panel (b) is a biofilm formed in urine containing 1.5 mg/ml of citrate with a nucleation pH of 8.3. The biofilm is sparse and is composed of micro-colonies of cells with no signs of crystalline material (from Stickler and Morgan (2006), with permission).

MacLeod and Stickler (2007) examined the associations between organisms in the biofilm communities on these catheters. It became clear that some organisms, e.g. Providencia stuartii or Klebsiella pneumoniae, were commonly present with P. mirabilis in the catheter biofilm communities. On the other hand, when Enterobacter cloacae or Morganella morganii were colonizing catheters, P. mirabilis was rarely present in the biofilm. The hypothesis that the absence of P. mirabilis from some biofilm communities could be due to its active exclusion by other species was then examined experimentally in laboratory models. It was found that when other species

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have established themselves in the models and formed biofilms on the catheters, most were capable of slowing catheter encrustation by P. mirabilis. These effects, however, were transient and in all cases P. mirabilis was able to raise the urinary pH, colonize the biofilms of the other species, induce crystal formation and block the catheters. Thus the presence of species such as E. cloacae or M. morganii in a catheterized urinary tract is not likely to prevent subsequent infection by P. mirabilis. The absence of P. mirabilis from catheter biofilms containing these species is thus probably because the patients concerned have just not been exposed to P. mirabilis. Infection with P. mirabilis whatever the pre-existing urinary flora is thus likely to lead to catheter encrustation and blockage. It is possible of course that more complex communities could have a more permanent inhibitory effect on P. mirabilis. Preventing colonization of the bladder urine of patients undergoing longterm catheterization is a tall order. Trautner and Darouiche (2004) suggested that it may be sensible to avoid the use of antibacterials in preventative strategies. They argued that the inevitable failure to suppress bacteriuria in these cases is related to the facility with which the numerous, diverse and frequently multi-resistant organisms from the bowel flora can gain access to the catheterized tract. The resistance problem is also compounded by the ease with which exchange of drug resistance genes occurs between the mixed populations of bacteria in the catheter biofilm. In view of these factors, they concluded that trying to rid the catheterized bladder of all microbial flora is futile and merely promotes the selection of resistant organisms. As an alternative approach they suggested that colonization of the catheterized urinary tract by benign bacteria might prevent or interfere with colonization by pathogenic species. It would be of great interest if such an interference strategy could be found to be effective against P. mirabilis.

7.7

Urease inhibitors

Clinical studies have shown that the oral administration of the urease inhibitor acetohydroxamic acid can reduce the extent of infectionassociated kidney and bladder stone formation (Griffith et al., 1978) and catheter encrustation (Burns and Gautier, 1984). Unfortunately, later studies reported that the acetohydroxamic acid treatment caused intolerable side effects and its use was abandoned (Williams et al., 1984; Griffith et al., 1991). Morris and Stickler (1998) used laboratory models to examine the effect of the urease inhibitors acetohydroxamic acid and fluorofamide on P. mirabilis-induced catheter encrustation. They demonstrated that when models were supplied with urine supplemented with acetohydroxamic acid (1.0 mg/ml) or fluorofamide (1.0 μg/ml) the rise in pH was controlled and significantly less deposition of calcium and magnesium salts compared with controls occurred on the catheters. It was suggested that it might be feasible

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to introduce these compounds directly into the bladder in order to avoid any systemic toxicity. Unfortunately, subsequent attempts to deliver either of these compounds through the retention balloon of all-silicone catheters proved unsuccessful (Williams, 2006). Resveratrol is a naturally occurring phytoalexin, one of a group of compounds that plants produce in response to attack by pathogens. It has antiinflammatory and anti-oxidant activity and recently has been shown to inhibit swarming and the production of urease by P. mirabilis (Wang et al., 2006). The presence of this compound in red wine opens up some interesting possibilities for the control of catheter encrustation! As a first step it would be well worth determining the concentration of resveratrol that would be required to prevent alkaline conditions developing in P. mirabilis-infected urine.

7.8

Catheter design

In his editorial in the New England Journal of Medicine, Calvin Kunin (1988) posed the rhetorical question ‘Can we build a better catheter?’. He expressed his disappointment that at a time when substantial technological advances were common in many other medical fields, it was still not possible to perform the relatively simple task of draining urine from a debilitated bladder without producing infections and their associated complications. He commented that catheter manufacturers have been reluctant to invest in research and development and that the catheters in use today were essentially the same as those introduced in the 1930s. Little has changed since this call for action. The catheters given to patients today with their roughly engineered irregular surfaces and their narrow central channels are readily colonized by bacteria and blocked by crystalline bacterial biofilms. Currently catheters are not costly, but the eventual price we have to pay for their use is enormous. There are fundamental design problems with Foley catheters that induce vulnerability to infection of the urinary tract. The very presence of the indwelling catheter undermines the bladder’s chief mechanical defence against infection. The cyclic filling and emptying of the bladder, which normally ensures that any bacteria contaminating the urine or trying to ascend the urethra are washed out, is compromised. This dynamic, regularly flushed system is replaced by one in which the presence of the eye-holes above the catheter balloon means that there is a stagnant sump of urine in the bladder. The volume of the residual urine is commonly around 100 ml (Garcia et al., 2006), from which there is a more or less continuous dribble of urine down the catheter. This establishes a continuous flow system into a reservoir of nutrient medium in which bacterial communities can flourish. The problems are compounded because the balloon and tip of the catheter can erode and damage the smooth mucosal surface of the bladder. The pressure exerted

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by the catheter on the walls of the urethra attenuates the blood supply and can block the lubricating peri-urethral glands. The stressed peri-urethral surface and the outer surface of the catheter provide convenient substrates for bacterial colonization and ascending migration into the bladder. In addition, the open channel of the catheter and drainage tube also offers easy access directly into the vulnerable bladder for organisms that have contaminated the drainage system. The Foley catheter thus constitutes a constant threat to the integrity of the bladder and upper urinary tract. As Kunin pointed out in his devastating critique of the catheter, the challenge is to produce a device that allows the tract to retain the normal physiological and mechanical characteristics of its voiding system. He suggested that this will require the design of a thinwalled, continuously lubricated, collapsible catheter so as to restore the integrity of the urethra. The retention balloon needs to be replaced by a device that does not result in a residual sump of urine and will permit regular intermittent complete washing out of the bladder. Such catheters would of course be both more difficult to manufacture and more expensive, but surely would be well worth the investment. It is possible to reinstate the periodic filling and emptying of the catheterized bladder by replacing the catheter drainage tube and bag with a simple manual valve (Fader et al., 1997; German et al., 1997). When the valve is closed the bladder fills with urine; regular opening of the valve then allows the periodic release of urine from the bladder. It has been suggested that this helps to maintain bladder tone and capacity. In addition, the use of the valve-might decrease bladder mucosa irritation, since periodic filling of the bladder will lift the catheter away from the bladder wall. It has also been claimed that the valve-regulated periodic flushing of the catheter might decrease infection and catheter encrustation (Addison, 1999; Doherty, 1999). Experiments in laboratory models of the catheterized bladder fitted with manual valves have demonstrated that opening the valve to drain urine from the bladder at 2 h intervals four times during the day, significantly delayed the blockage of the catheter by P. mirabilis crystalline biofilm. An even greater beneficial effect was recorded when a timercontrolled valve was used to achieve release of urine at 4 h intervals throughout the day and night (Sabbuba et al., 2005). As manual valves present difficulties to many people with poor manual dexterity, electronically controlled automatic valves are being developed. One such compact, reliable prototype having a suitably low power consumption shows considerable promise (Lee et al., 2007). Other valves have been developed that respond to pressure changes in the bladder (e.g. The Urocycler, UroSolutions Inc., Orlando, Florida). Even with the application of a valve, the bladder is not completely emptied. The presence of the balloon still ensures the residual sump of

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urine. Alternatives to the balloon, such as expandable and retractable funnels designed to allow the catheterized bladder to empty completely, are being developed but mechanisms for opening the retention devices once the catheter has been located in the bladder and the subsequent retraction of the device for catheter removal need to be perfected. An innovation that could greatly reduce the amount of residual urine in the bladder is the advanced catheter system (ACS) invented by Professor Roger Feneley of The Biomed Centre at the Bristol Urological Institute, UK. This consists of silicone tubing that is inserted suprapubically and passes through the bladder and exits through the urethra. A series of drainage holes in the bladder section facilitates the drainage of urine ensuring minimal residual urine. The catheter can be drained from either the suprapubic or urethral end. Debris accumulating in the bladder, as a result of P. mirabilis infection for example, can also be flushed out by either continuous irrigation or by temporary clamping of the tubing and the instillation of appropriate solutions. In the case of patients with prostatic hypertrophy, it is the prostatic urethra that has to be kept open. Intraurethral devices have been developed as alternatives to catheters since the full length of the catheter is not required. Intraprostatic metallic spirals such as the Prostakath or Urospiral have been used but blockage by encrustation remains a problem (Braf et al., 1996). More recently, a self-expandable, biodegradable spiral stent, the Spiroflow, manufactured from an l-lactide-glycolic acid co-polymer has been described and has been reported to be less vulnerable than the metallic devices to encrustation in urine (Laaksovirta et al., 2003). It should be pointed out, however, that these tests were performed in vitro, the devices only being exposed to uninfected urine. Another novel intraurethral device is the In-FlowTM valve–pump catheter. This has been designed for women with hypocontractile or acontractile bladders, as an alternative to indwelling or clean intermittent catheterization. It is a short silicone catheter containing an internal valve and pump mechanism operated by an external remote control unit. The device is introduced into the urethra by a disposable inserter. It is fixed in position by flexible silicone fins that open like flower petals at the level of the bladder neck and by a flange at the external urethral meatus. The device is designed to be replaced routinely at 28 day intervals. This is achieved by manipulation of the flange to collapse the retention fins. The valve and pump mechanism is a magnet located in the core of the catheter. To operate the valve, an activator unit powered by two 3 volt lithium batteries is held at the pubic area and the ‘on’ button is depressed. The valve opens and miniature rotor spins at 10 000 rpm, generating a urine flow rate of 10–12 ml/s. On release of the button, the magnet automatically counter spins and the valve is closed. A multi-centre study of this device found that

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although most (169 of 273) patients withdrew from the programme because of discomfort and leakage, those that completed a 16 week trial period were reported to have improved their quality of life scores (Chen et al., 2005). It was also encouraging that the mean post-voidal residual volume of urine in these patients was only 16 ml. Little detail was given on the bacteriology, except that 50% of those that completed the study had asymptomatic bacteriuria. It would have been interesting to know how the In-FlowTM device performed in patients infected with P. mirabilis. A catheter with an intraurethral section that was so soft and compliant that it conformed to the shape of the urethra was described some 20 years ago (Brocklehurst et al., 1988). It was reported that not only did patients find these catheters more comfortable but also significantly fewer (P < 0.001) of the conformable catheters had intraluminal obstruction with struvite compared with silicone-coated latex control catheters. Despite these encouraging results this catheter has not been developed further. Another early idea that has not been taken up is that of self-lubricating catheters. Kunin and Finkelberg (1971) suggested that because in many cases the urethra becomes dry, irritated and inflamed on long-term catheterization, it would be beneficial if the catheter was repeatedly lubricated. They developed a prototype in which lubricant was extruded intermittently through an additional channel within the catheter. A controlled trial showed a significant reduction in bacteriuria in females who received daily methyl cellulose or polymyxin B lubricant compared with those who had catheters that were not lubricated. They concluded that if a less cumbersome delivery system could be developed it might have considerable potential. The recent ingenious development of self-lubricating silicones is thus of great interest (Malcom et al., 2004). These silicones were produced by the condensation cure of a poly(dimethylsiloxane) with a special cross-linker that generates oleyl alcohol as a product in the polymerization process. Oleyl alcohol is lubricious and is formed throughout the elastomer matrix. Slow diffusion of oleyl alcohol through the matrix and its persistence at the surface of the material results in a highly lubricious silicone biomaterial. These polymers were also reported to be capable of delivering antibacterial agents to the surrounding medium. Metronidazole was incorporated as a model drug into these self-lubricating elastomers and it was found that the rate of metronidazole flux from the polymer could be manipulated by varying the concentration of the cross-linker. These self-lubricating silicones loaded with an agent targeted against P. mirabilis could well find applications in the prevention of catheter encrustation and blockage. If the medical device industry cannot be persuaded to exploit some of these ideas, then at least we should encourage them to make relatively simple improvements in the current Foley catheter design and manufacturing standards. The narrow internal diameters of the current catheters

Proteus mirabilis biofilm formation and catheter design (a)

(b)

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(c)

2 mm

7.9 Photographs showing cross-sections of unused catheters: (a) a silicone-coated latex catheter; (b) a hydrogel-coated latex catheter; (c) an all-silicone catheter (from Stickler et al. (2003a), with permission).

(Fig. 7.9) induce acute vulnerability to blockage. Alternatives to latex and silicone as the base materials for manufacture might be considered. For example, prototypes have been made from a nylon braid encased in a polyurethane matrix which has excellent flow properties and resistance to kinking; they have substantially thinner walls giving much larger internal diameters that are not easily occluded (Lawrence and Turner, 2006). Surely there are also better designs and manufacturing processes for producing eyelets with smoother surfaces, larger orifices and improved flow characteristics.

7.9

Future trends

It is difficult to understand why at a time when amazing technological advances are being made in other areas of medicine, we are still unable to drain urine from a debilitated bladder without producing infection and a whole range of associated complications. The morbidity caused to patients and the costs to health authorities resulting from the use of the long-term indwelling bladder catheter are surely no longer acceptable. It has to be realized that attempts to prevent encrustation merely by impregnating current catheters with antibacterials or by coating them with surfaces that inhibit bacterial adhesion are unlikely to be successful in the long run. These approaches show a lack of understanding of how the crystalline biofilms develop and underestimate the ability of microbes such as P. mirabilis to evolve in such a way as to overcome attempts at control using chemical agents. Advances must be made in catheter design. Better eye-holes and wider internal diameters should be simple to achieve but fundamental improvements will demand the design of devices that reinstate the emptying and filling of the bladder, do not produce a residual sump of urine in the bladder, and do not irritate and inflame urethral and bladder epithelia.

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Alternative retention devices that can be easily deployed and withdrawn, and automatic valves that ensure filling and then complete emptying of the bladder are required. It is unlikely that a singly device is going to provide the universal solution to the problems of bladder management for all patients. A series of devices tailored to the specific needs of various groups of patients with different disabilities and requirements will have to be developed. In the meantime, while P. mirabilis remains such a problem it is important to keep calcium and magnesium phosphates in solution by preventing the urinary pH rising above its pHn. P. mirabilis is an ingenious organism but even it can do little about the laws of physical chemistry that govern the solubility of phosphates. A simple device to determine the pHn of urine would be useful in patient management to signal circumstances where steps should be taken to increase fluid intake. Calvin Kunin had the foresight in 1988 to realize that in the face of the reluctance of the catheter industry to invest in research and development or to change their manufacturing processes, a sea change is required in the way we approach research in this field. To outline the necessary changes, I can do no better than quote from the conclusion of his editorial in the New England Journal of Medicine of 1988: We must therefore consider initiating a national program that will combine the efforts of government, investment by industry, and the talent in academic institutions to develop better indwelling catheters and bring them to clinical trial. The catheters currently available are inexpensive instruments, but the eventual price we have to pay is great indeed. Catheters of the future may well be more expensive, but should be well worth the investment.

7.10

Conclusions

Infection with P. mirabilis exposes the many design faults of the currently available Foley catheters with potentially disastrous consequences for patient care and substantial financial implications for health authorities. It is clear that higher standards are required in the design and the manufacture of these devices. The catheters available today with their roughly engineered surfaces and thick walls and narrow central channels are extremely vulnerable to blockage by crystalline P. mirabilis biofilms. There is plenty of scope for improvement and there are many interesting ideas in the literature. Building a better catheter is certainly possible. The morbidity associated with the infections that result from the current devices undermines the quality of life and threatens the health of so many people and is no longer acceptable. Catheter manufacturers should take up the challenge.

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7.11

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Sources of further information and advice

I can recommend the book edited by Mobley and Warren Urinary Tract Infections Molecular Pathogenesis and Clinical Management ASM Press, Washington, 1996 as a good source of information about the basic biology, epidemiology and pathogenicity of Proteus mirabilis. The recent extensive review by Jacobsen et al. 2008, Clinical Microbiology Reviews, will update the student on the nature of the virulence factors of P. mirabilis and their involvement in catheter-associated infections. Calvin Kunin’s book Urinary Tract Infections: Detection Prevention and Management is a gold mine of valuable basic information and has a particularly useful chapter on catheter care. Disinfection and Decontamination: Principles, Applications and other Issues, edited by Gurusamy Manivannan, CRC Press, Boca Raton, 2007 has a chapter that gives a detailed review of attempts to develop antimicrobial catheters. Kunin’s editorial ‘Can we build a better urinary catheter?’ in the New England Journal of Medicine (1988), 319: 365–366, should be compulsory reading for those interested in improving the care of the many people disabled by loss of bladder function. The Biomed Centre, established by Professor Roger Feneley in the Bristol Urological Institute at Southmead Hospital Bristol, UK, was set up especially to deal with the issues of long-term care of patients with disabled bladders. The Healthcare Technology Cooperative at the Biomed Centre with its clinical, laboratory and training facilities, and through its collaborative network of clinicians, engineers, polymer chemists and microbiologists, based in industry, academia and other clinical centres, aims to accelerate the development and provision of new devices, technologies and procedures designed to improve the health and quality of life of patients with severe urinary incontinence. The E-mail address is [email protected] and the website is www.biomedhtc.org.uk.

7.12

References

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stickler d j and sabbuba n a (2008), ‘Antimicrobial catheters’, in Manivannan G (Ed.), Disinfection and Decontamination, Boca Raton, CRC Press, pp. 415–457. stickler d j, ganderton l, king j, nettleton j and winters c (1993), ‘Proteus mirabilis biofilms and the encrustation of urethral catheters’, Urol Res, 21, 407–411. stickler d j, morris n, moreno m c and sabbuba n a (1998), ‘Studies on the formation of crystalline bacterial biofilms on urethral catheters’, Eur J Clin microbial Infect Dis, 17, 649–652. stickler d j, young r, jones g, sabbuba n s and morris n s (2003a), ‘Why are Foley catheters so vulnerable to encrustation and blockage by crystalline bacterial biofilm?’, Urol Res, 31, 306–311. stickler d j, jones g l and russell a d (2003b), ‘Control of encrustation and blockage of Foley catheters’, Lancet, 361, 1435–1437. stickler d j, lear j c, morris n s, macleod s m, downer a, cadd d h and feast w j (2006a), ‘Observations on the adherence of Proteus mirabilis onto polymer surfaces’, J Appl Microbiol, 100, 1028–1033. stickler d j, jones s m, adusei g o and waters m g (2006b) ‘A sensor to detect the early stages in the development of crystalline Proteus mirabilis biofilm on indwelling bladder catheters’, J Clin Microbiol, 44, 1540–1542. stickler d j, jones s m, adusei g o, waters m g, cloete j, mathur s and feneley r c (2006c), ‘A clinical assessment of the performance of a sensor to detect crystalline biofilm formation on indwelling bladder catheters’, Br J Urol Int, 98, 1244–1249. suller m t, anthony v j, mathur s, feneley r c, greenman j and stickler d j (2005), ‘Factors modulating the pH at which calcium and magnesium phosphates precipitate in human urine’, Urol Res, 33, 254–260. trautner b w and darouiche r o (2004), ‘Catheter-associated infection: pathogenesis prevents prevention’, Arch Intern Med, 164, 842–850. wang w b, lai h c, hseuh p r, chiou r y y, lin s b and liaw s j (2006), ‘Inhibition of swarming and virulence factor expression in Proteus mirabilis by resveratrol’, J Med Microbiol, 55, 1313–1321. warren j w, tenney j h, hoopes j m, muncie h l and anthony w c (1982), ‘A prospective microbiologic study of bacteriuria in patients with chronic indwelling urethral catheters’, J Infect Dis, 146, 719–723. williams f d and schwarzhoff r h (1987), ‘Nature of the swarming phenomenon in Proteus’, Ann Rev Microbiol, 32, 101–122. williams g j (2006), ‘The use of antimicrobial agents to control the development of crystalline Proteus mirabilis biofilms on urinary catheters’, PhD thesis, Cardiff University, Wales, UK. williams g j and stickler d j (2007), ‘Some observations on the diffusion of antimicrobial agents through the retention balloons of Foley catheters’, J Urol, 178, 697–701. williams j j, rodman j s and peterson c m (1984), ‘A randomized double-blind study of acetohydroxamic acid in struvite nephrolithiasis’, N Engl J Med, 311, 760–764.

8 Self-lubricating catheter materials A. D. W O O L F S O N, R . K . M A L C O L M, S. P. G O R M A N and S. D. M C C U L L AG H, Queen’s University Belfast, UK

Abstract: This chapter discusses the problem of lack of lubricity in common biomaterials, notably silicone elastomers, used in the production of urinary catheters. The chapter first considers the basis of silicone chemistry for elastomer production, leading to a description of the chemistry of novel, self-lubricating silicone biomaterials. Finally, the characteristics of these novel materials are described with respect to lubricity, mechanical performance and biocompatibility aspects. Key words: silicone, elastomer, lubricity, urinary catheter, biocompatibility.

8.1

Introduction

The availability of medical devices for temporary or permanent implantation constitutes a major advance in modern medicine. Unfortunately, as the use of such devices becomes increasingly common, so does the incidence of complications associated with their use – such as infection, encrustation and flow obstruction (Gorman and Jones, 2003). Urinary tract devices, for example, present a particular problem in respect of blockage and the resulting device failure owing to flow obstruction (Kunin, 1989). Urethral catheters may remain indwelling for weeks to years. Long-term catheterisation (LTC) of the bladder is considered a method of last resort with its use occurring commonly in the elderly as well as those with chronic urinary incontinence or urethral obstruction. For either short- or long-term catheter usage, the infection rate is about 5% per day (Nicolle, 2005). Device-related infection is a major life-threatening problem with all types of medical device owing to the formation of a microbial biofilm on the biomaterial. When this occurs, antibiotic therapy is unable to eradicate the infection and invasive surgery may be necessary to remove and replace the device. Improvements in medical device biomaterials are urgently required to prevent these problems. A further issue that presents difficulty for patients is the inherent lack of lubricity associated with all common catheter biomaterials (Jones et al., 191

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2004). Thus, the device may be difficult to insert, although this problem can be overcome to some extent by use of a lubricating gel. More importantly, for an indwelling silicone catheter, typically in place for up to 3 months, the device is difficult to remove, the lack of lubricity often causing epithelial tissue damage upon removal. Mechanical irritation of the urethra may result in mucosal abrasion, subsequently leading to submucosal inflammation and infection (Khoury et al., 1991). Silicone elastomer is a common rubbery biomaterial presently used in the production of urinary catheters and similar devices (Habal, 1984). Silicones are inorganic polysiloxane (silicone) compounds consisting of a repeating silicon–oxygen backbone unit, of which the most common is polydimethylsiloxane (PDMS). PDMS in crosslinked, elastomeric (rubber) form has been used for many years in a large range of medical devices such as urethral or vascular catheters, and as implantable drug delivery devices. Silicone elastomeric devices combine ease of fabrication, processability and biocompatibility, and are relatively inexpensive to manufacture. However, in common with all current biomaterials, silicone is prone to biofilm formation and, in the case of urinary catheters, to encrustation with complex inorganic salts (Stickler et al., 1998). In addition, silicone lacks lubricity (Woolfson et al., 2003). There are many different crosslinking techniques for the manufacture of silicone elastomers (Braley, 1970). However, in the medical devices industry only two systems are routinely employed: the platinum catalysed hydrosilation reaction or the alkoxy condensation reaction system. In the condensation reaction, a tri- or tetra-functional alkoxysilane is reacted with an α, ω-hydroxy terminated PDMS polymer producing the crosslinked elastomer as well as the condensation alcohol, usually propanol. The propanol reaction product then permeates through the elastomer to the surface, where it diffuses into the surrounding medium. This well-known crosslinking system for the production of silicone elastomers offers an intriguing possibility for the design of self-lubricating silicone elastomers (Woolfson et al., 2003). Thus, this chapter introduces the basis of silicone chemistry for elastomer production, leading to the chemistry of novel, self-lubricating silicone biomaterials and a consideration of their performance characteristics.

8.2

Silicone chemistry

Organosilicons and alkoxysilanes are synthetic compounds primarily manufactured from chlorosilanes. In 1823, Berzelius synthesised tetrachlorosilane, SiCl4 (Colvin, 1981) and from this compound, Friedel and Crafts in 1863 manufactured the first organosilane, tetraethylsilane, and Ebelman in 1846 synthesised the first alkoxysilane, silicon isoamyloxide (Bradley et al., 1978). In the late nineteenth and early twentieth centuries Kipping

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investigated the properties of a range of organosilanes and, coupled with the Grignard reaction, manufactured the first polymeric compound. After identifying the silicon–oxygen linkage and (incorrectly) presuming the presence of an Si=O bond, in analogy with ketones, he conceived the term ‘silicones’. Nonetheless, he concluded that ‘the few (organic derivatives of silicon) which are known are very limited in their reactions, the prospects of any immediate and important advance in this section of organic chemistry does not seem to be hopeful’ (Colvin, 1981). However, within a few years, Corning Glass Works of America manufactured organosilanes to prepare the first silicone resin. Continuing research into the formulation of the silicone polymers led to the production of the first silicone rubber in 1945 which, although similar in physical properties to natural rubbers, differed significantly in its chemical structure (Levin, 1958). Pharmaceutical applications for silicone materials began in the 1950s and the first implantation was credited to De Niola in 1959, using silicone rubber tubing as a replacement for a urethra (Habal, 1984). The first silicone elastomers of the room temperature vulcanisation (RTV) type, that could be cured in situ without heat, were introduced in the 1960s. Subsequently, silicones have been used in diverse medical applications such as implants, drug delivery devices and catheters.

8.2.1 Silicone nomenclature Silicone is the generic name for a class of polymers based on the pattern of a repeating silicon–oxygen backbone, the most common of which, PDMS, consists of the repeating unit, -[Si(CH3)2-O]- (Fig. 8.1). By replacing the methyl units with different groups – for example phenyl, ethyl and vinyl groups – silicones with different physical and chemical properties (Arkles, 1983) can be produced (Table 8.1). The most commonly used polymer nomenclature for silicones, the General Electric siloxane notation, utilises the letters M, D, T and Q to represent mono-, di-, tri- and quadri-functional monomer units respectively (Fig. 8.2). Unprimed letters represent methyl substituents and primed letters represent other substituents; for example (CH3)(C6H5)SiO is represented as D′ whereas (CH3)2SiO is represented by D (Ratner, 2004).

8.2.2 Crosslinking (vulcanisation) of silicones for elastomer production Crosslinking or vulcanisation of a polymer involves the formation of crosslinks between separate polymer chains in order to achieve a threedimensional network. Typically, these crosslinks are separated from each

194

Biomaterials and tissue engineering in urology CH3 H3C

Si

CH3 O

CH3

Si

CH3 O

CH3

n

Si

CH3

CH3

8.1 Chemical structure of polydimethylsiloxane.

R4O

R2O

(Q)

(T)

OR1

CH3

Si

OR2

R3O

Si

OR3

OR1

(D) CH3

(M)

Si

OR2

CH3 OR1

R1O

CH3

Si

CH3

CH3

8.2 General Electric siloxane notation.

Table 8.1 Commonly substituted groups with their applications Structure

Name

Application

CH3C6H5-

Methyl Phenyl

CF3CH2CH2HHO-

Trifluoropropyl Hydride Silanol

CH2=CH-

Vinyl

Basic substitution found in all silicones Increases modulus, thermal and ultraviolet stability Increases stability and solvent resistance Introduces crosslinking site Crosslinking site for condensation and metal catalysed crosslinking Increases peroxide reactivity

other by chain segments ranging from between 100 and 400 repeating monomer units (Polmanteer, 1981). In addition to conventional covalently bonded systems, polymers have been crosslinked by dipole–dipole and Van der Waals forces. Furthermore, temporary crosslinks can be formed by trapped chain entanglements. The main reaction mechanisms used for the vulcanisation of medical silicones are condensation or addition cures. Addition cures are based upon the peroxide cure or the hydrosilation reaction, with the latter system used

Self-lubricating catheter materials SiH

+

(Catalyst)

H2C

C H

H

H

Si

195

Si

Si H

H

8.3 Formation of an ethylene bridge during hydrosilation.

for medical grade materials. Hydrosilation crosslinking of silicone polymers involves the addition of a silicon–hydrogen bond to a carbon–carbon group in the presence of a catalyst, usually chloroplatinic acid, covalently linking the siloxane chains with short-chain oligomers (Fig. 8.3). Generally, a shortchained silicone molecule having several silicon hydride groups (Si-H), often referred to as the crosslinker (Speier et al., 1957), is mixed with a high molecular weight silicone polymer having several vinyl groups attached. A feature of this curing system is that both the polymer and the crosslinker can have multiple functional groups, which can be located either on pendant or terminal positions on the siloxane chain. By varying the mixture of polymers and the chain lengths, the density and the degree of crosslinking can be controlled to produce elastomers of both low and high consistency. The reaction mechanism for hydrosilation is shown in Fig. 8.4. The platinum complex, H[(C3H6)PtCl3], catalyses the hydrosilation process via the Chalk and Harrod method (Chalk and Harrod, 1965). The mechanism is initiated by the weakening of the silicone vinyl group by the formation of a complex with the catalyst. Next, the catalyst activates the hydrosilane by the cleavage of the Si–H bond. The platinum metal atom is thought to provide a site in which the two species are in the proper proximity and stereochemical configuration for addition to occur, producing an ethylene linkage between the two (Chalk and Harrod, 1965). The catalyst is then regenerated. Condensation cures are characterised by the formation of a chemical bond between two compounds, with the ensuing elimination of a small compound such as water or methanol. There are several crosslinking compounds used for this purpose and they are based upon the concept of a tri- or tetra-functionalised crosslinking reagent. The most common crosslinking reagents are the acyloxy, amine, oxime, alkoxy and silicon hydride systems. In general, the amines and oximes react the fastest followed by the acyloxy and then the alkoxy systems (Arkles, 1983). The alkoxy system (Fig. 8.5) is the most common technique for crosslinking RTV condensationcured medical silicones. The reaction is often aided by adding an organotin catalyst in medical applications. The mechanism of the reaction of organotin catalysts, specifically dialkyltin dicarboxylates, was investigated by Van der

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H +

Pt

HC

SiH

Si

Pt

Si

+

HC

Si Pt

Pt

8.4 Mechanism of platinum catalysed hydrosilation.

OCH 2 CH2 CH 3

CH 3 HO

Si

O

n

+

H

CH 3 CH2 CH 2O

Si

OCH 2 CH 2 CH 3

OCH 2 CH2 CH 3

CH 3 Hydroxy terminated α, ω polydimethylsiloxane

Tetrapropoxysilane Catalyst * O

H3 C CH 3 *

O

Si

Si O

n

O

CH 3 H3 C

Si

CH3

n O

O Si O

CH3 Si

O

n*

CH3 CH3

n

* Crosslinked polydimethylsiloxane

+ CH 3CH 2 CH2 OH Propanol

8.5 Mechanism of alkoxy crosslinking for the synthesis of condensation-cure silicone elastomer.

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197

Weijj (1980). The reaction proceeds when moisture present in the formulation partially hydrolyses the catalyst forming an active tin hydroxide species. This species in turn reacts with an alkoxysilane, for example tetrapropoxysilane (TPOS) shown in Fig. 8.5, to produce an organotin silanoate. The organotin silanoate reacts with the α, ω-hydroxy terminated PDMS, giving rise to the formation of the crosslinked network via the silanolysis of the Sn–O–Si bond. The organotin hydroxide is then regenerated, allowing further reactions to occur. The resulting silicone elastomer is often referred to as an RTV silicone, since the reaction will occur at room temperature. Typically, though, a temperature between 60 and 100 °C is used in order to have a fast cure. The condensation by-product formed, a low molecular weight alcohol, diffuses to the surface and either evaporates or, if immersed in a solvent, is removed by dissolution. Normally, this is a useful feature since it removes a reaction impurity from the final product. However, this latter reaction gives rise to the application of novel chemistry for the design of self-lubricating silicone biomaterials.

8.3

Self-lubricating silicone biomaterials

Conventional RTV silicone elastomers are produced according to the scheme described in Fig. 8.5, with a volatile, low molecular weight alcohol as a by-product. The alkoxysilane ‘four-armed’ crosslinker used to link on to the hydroxyl terminal groups at each end of PDMS is usually TPOS. A new crosslinker, again with four ‘arms’ or crosslinking points, can be made by the reaction of TPOS with 1-octyl-2-dodecanol in a 1 : 4 molar ratio transetherification reaction that produces non-volatile tetraoctyldodecoxysilane (TODDOS, Fig. 8.6), a viscous, oily liquid, and a volatile propanol by-product. Similarly, a range of non-volatile, higher silane crosslinkers OC3H7 H7C3O

Si

OC3H7

OC3H7 TPOS OCH3(CH2)9CH[(CH2)7CH3]CH2OH HOCH2[(CH2)7CH3]CH(CH2)9CH3O

Si

OCH3(CH2)9CH[(CH2)7CH3]CH2OH

OCH3(CH2)9CH[(CH2)7CH3)]CH 2OH TODDOS

8.6 The structures of conventional (TPOS) and novel (TODDOS) silicone crosslinkers.

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Table 8.2 Tetra-n-alkoxysilane crosslinkers for RTV silicones Tetra-n-alkoxysilane crosslinker

Molecular structure

Tetrapropoxysilane (standard)

Abbreviation TPOS

OC3H7 H7C3O

Si

OC3H7

OC3H7

Tetrabutoxysilane

TBOS

OC4H9 H9C4O

Si

OC4H9

OC4H9

Tetrahexoxysilane

THOS

OC6H13 H13C6O

Si

OC6H13

OC6H13

Tetraoctoxysilane

TOOS

OC8H17 H17C8O

Si

OC8H17

OC8H17

Tetradecoxysilane

TDOS

OC10H21 H21C10O

Si

OC10H21

OC10H21

Tetrastearoxysilane

TSOS

OC18H37 H37C18O

Si

OC18H37

OC18H37

Tetraoleyloxysilane

OC9H18 H18C9

H18C9O

Si

OC9H18

OC9H18

TOLOS

C9H18

C9H18

C9H18

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199

can be prepared (Woolfson et al. 2003), for example tetraoleyloxysilane (TOLOS), as shown in Table 8.2. When a higher silane crosslinker is substituted, wholly or partly, for TPOS in the reaction scheme shown in Fig. 8.5, the crosslinking reaction proceeds as shown, with each of the four crosslinker arms forming a crosslinking site as part of the resulting threedimensional elastomeric silicone network. The by-product of the reaction in these cases is not propanol, however, but the parent, non-volatile higher alcohol used to form the silane crosslinker. Thus, for example, the use of TOLOS as a crosslinker results in the formation of a clear, oily liquid, oleyl alcohol, within the elastomer matrix (Woolfson et al. 2003). This is the principle that gives rise to a self-lubricating silicone elastomer biomaterial. The oily higher alcohol ‘blooms’ to the surface following crosslinking to produce a thin, oily film on the elastomer surface that is readily replenished from the reservoir of oil within the elastomer matrix, allowing for sustained selflubrication of the silicone surface. Importantly, this provides for lubricity of a device on both insertion and removal.

8.4

Performance characteristics of self-lubricating silicone biomaterials

8.4.1 Determination of the coefficients of friction The use of hydrogel or similar coatings with low coefficients of friction is a common method of addressing the clinical difficulties associated with insertion and removal of medical devices (Tunney and Gorman, 2002). However, a biomaterial that is always highly lubricious, yet does not require the use of coating technology, would clearly be advantageous. The coefficients of both static and dynamic friction were measured for a range of TODDOSderived lubricious silicones using the apparatus shown in Fig. 8.7. The silicone biomaterials were prepared in flat sheet form. At the start of each experiment the stainless steel sled (base area 2.375 × 10−3 m2) is placed lightly on the surface of the test material such that the

Moving upper clamp Waxed cotton thread Weighted stainless steel sled Test material

8.7 Experimental design for determination of the coefficient of friction.

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Table 8.3 Static and dynamic coefficients of friction for lubricious (TODDOSderived) RTV silicone elastomers and for standard RTV silicone. Each figure is the mean ± one standard deviation of eight replicates

Crosslinker composition

Coefficient of static friction, μs

80% TODDOS/20% TPOS 60% TODDOS/40% TPOS 50% TODDOS/50% TPOS 100% TPOS

0.084 0.080 0.085 0.531

± ± ± ±

0.005 0.006 0.005 0.046

Coefficient of dynamic friction, μd 0.075 0.079 0.063 0.312

± ± ± ±

0.018 0.011 0.016 0.042

thread attaching it to the upper clamp (via the pulley) is just taut. The crosshead is then moved vertically upwards through a distance of 50 mm at a constant speed of 4 mm s−1. This causes the sled to move horizontally across the surface of the silicone elastomer towards the pulley through the same distance and at the same speed. The forces required to produce movement are recorded and used to calculate the static and dynamic coefficients of friction, from the force required to initiate movement or the mean force recorded during movement respectively. The coefficients of friction are calculated by dividing the force required to produce movement by the downward force owing to the weight of the sled. Friction tests have been performed for a range of lubricious formulations and for standard RTV silicone crosslinked with TPOS. For all lubricious silicone formulations, a smooth sliding movement was observed using a sled with a weight of 1.1 kg. However, for the standard RTV silicone the sled did not slide but toppled over under the pulling force. Therefore, the sled weight was reduced to 0.234 kg for this sample. The frictional coefficients for a range of TODDOS silicones partly crosslinked with TPOS and TODDOS in percentage crosslinker values from 80 to 50% TODDOS are shown in Table 8.3. The significant reductions in friction, in both static and dynamic modes, compared with a parent RTV silicone, i.e. crosslinked with 100% TPOS, are quite apparent. The variation in the amount of TODDOS compared with TPOS crosslinker has no discernible effect on the frictional coefficients, as would be expected since these depend only on a surface covering of oily exudate. The amount of TODDOS used as a percentage of the total crosslinker concentration will only affect the amount of total 1-octyl-2-dodecanol (ODD) formed within the elastomer and thus the duration of the lubricious effect.

8.4.2 Determination of mechanical strength It is apparent from the reaction scheme in Fig. 8.5 that the chemical structure of the silicone biomaterial produced by using either TOLOS for TPOS

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201

Table 8.4 Tensile properties of lubricious (TODDOS-derived) and standard RTV silicone elastomers measured 1 day after manufacture. Each result is the mean ± one standard deviation of at least five replicates

Crosslinker composition

Tensile strength at break (MPa)

Strain at break

80% TODDOS/20% TPOS 60% TODDOS/40% TPOS 50% TODDOS/50% TPOS 100% TPOS

0.580 ± 0.064 1.144 ± 0.105 1.415 ± 0.096 3.37 ± 0.19

2.405 1.473 1.659 1.49

± ± ± ±

0.177 0.116 0.144 0.09

as the crosslinking agent will be identical. However, the mechanical (tensile) properties may well differ, depending on the number of crosslinks that are formed. This, in turn, will depend on the efficiency of the crosslinking reaction. Tensile analysis of lubricious silicone elastomers can be performed using a TA-XT2 Texture Analyser (Jones et al., 1996) with a suitable sample clamping system. Elastomer test specimen dimensions were 10 mm (width) × 70 mm (length) × 1.6 mm (height). Specimens were firmly clamped (with the long axis vertical) such that there was a 50 mm length of material between the upper and lower clamps. Care was taken to ensure that there was no tension on the test specimen at the start of the determination. The lower clamp was static, while the crosshead was moved vertically upwards at a constant crosshead speed of 1.0 mm s−1. The sample tensile strength was calculated by dividing the load at break (Newtons) by the crosssectional area of the specimen (m2) and expressed in MPa. The strain at break was calculated by dividing the extension of the sample at break by its original length, since the force–time profile for all samples was essentially linear throughout. Therefore, the modulus of elasticity (Young’s modulus) was calculated simply by dividing the tensile strength at break by the strain at break and expressed in MPa. The data for mixed TODDOS/ TPOS crosslinked silicone RTV elastomers (Table 8.4) show that there is some reduction in tensile strength as a result, presumably, of less efficient crosslinking by the larger, and possibly more sterically hindered, TODDOS. The higher the amount of TPOS used in the crosslinking mixture, the less is the reduction in mechanical strength. Nevertheless, good quality elastomers were produced in all cases.

8.5

Bioactive lubricious silicones

Although inherent lubricity overcomes some of the problems associated with silicone medical devices, there remains the problem of microbial

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biofilm formation in vivo. The controlled delivery of an antimicrobial agent from the device is one means that may be used to overcome this problem. However, the drug release mechanisms operating in lubricious silicone biomaterials may be expected to differ significantly from conventional condensation-cure silicone elastomers (Malcolm et al., 2004). One solution to this problem may be to incorporate a suitable antimicrobial agent into the biomaterial, thus adding, in the case of a urinary tract device, a secondary drug delivery function to the primary mechanical function of the device. In order to demonstrate this possibility a model antibacterial agent, metronidazole, at drug loadings from 1.0 to 10.0% w/w can be incorporated into a lubricious silicone elastomer wholly or partly crosslinked with TOLOS (McCullagh et al., 2004). The release profile of the drug is conventionally determined under sink conditions, where the solubility in the release medium is sufficient to ensure that drug release from the sample is not impeded by a saturated receiving fluid. The release profile is important since it will dictate the duration of drug release, the amount of drug released instantaneously and cumulatively in use. Metronidazole was released from TOLOS lubricious silicone matrices according to t1/2 kinetics. Accordingly, straight-line cumulative release per unit area (Q) versus root time plots were obtained for all systems. The metronidazole flux rates are observed to increase with increasing TOLOS concentration and with increasing drug loading, the latter as predicted by the Higuchi equation (Ritger and Peppas, 1987), commonly used to model the release of substances from non-biodegradable, diffusion-controlled matrices under sink conditions. Completely replacing the standard TPOS crosslinker with TOLOS enhanced the metronidazole flux rate by between 52 and 65% depending on drug loading. Even the replacement of 20 molar percentage TPOS with TOLOS increases the flux rate by between 22 and 32%. It is therefore clear that in addition to the lubricious character afforded by the TOLOS silicone system, the oleyl alcohol by-product of the crosslinking reaction has also modified the permeation rate of metronidazole through the silicone elastomer. Given that the process of permeation in silicone elastomer systems is governed by solubility and diffusional factors (McCullagh et al., 2004), it is apparent that the oleyl alcohol is either enhancing release by providing increased solubility of the metronidazole in the system, or by decreasing the diffusional resistance in the system for the solubilised drug molecule, or through a combination of both mechanisms. Solubility studies (McCullagh et al., 2004) suggest that the enhanced release is attributed to the increased solubility of metronidazole in oleyl alcohol relative to that of the silicone elastomer. Enhanced release of a relatively water-soluble agent such as metronidazole from a hydrophobic, rather unfavourable environment such as silicone is a useful additional feature of the lubricious RTV silicones.

Self-lubricating catheter materials

8.6

203

Biomimetic lubricious silicones

The term ‘biomimetic’ in respect of a non-biological surface embraces properties that are substantially similar to human or animal epithelial tissue in vivo. Thus, human or animal epithelial body surfaces are protected by constant renewal through the production and shedding of mucous, which provides surface lubrication. Biomimetic surface properties, therefore, include: resistance to surface microbial growth and infection; resistance to surface deposition of solid material – such as, for example, the deposition of complex inorganic salts; and lubricity. In common with all other widely used biomaterials, silicone elastomer used in the production of urinary catheters is prone to biofilm formation and, in the case of silicone urinary devices, to encrustation with complex inorganic salts. If the exudate from lubricious silicones were to be removed constantly from the surface of the biomaterial in use, it may be argued that such a system is biomimetic. The incorporation of a suitable surface active agent in the preparation of lubricious silicone elastomers provides release to the elastomer surface such that the surface exudates are emulsified. Such emulsified surface exudates are then readily removed along with attached encrusting deposits and micro-organisms. The removal of attached material from the silicone elastomers of the proposed invention may be further optimised by application of a catheter maintenance solution to the elastomer surface in vivo. Such catheter maintenance solutions act on the emulsified surface exudate of the elastomer to allow cleaning of the surface and removal of unwanted attached material. The effect of incorporation of the surfactant benzalkonium chloride (1% w/w) in a TOLOS-derived lubricious silicone was determined against microbial biofilm and encrustation. Test samples of this material were then coated with bacteria and encrusted deposits prior to being subjected to 15 min contact with benzalkonium chloride (0.005% w/v at pH 4.5) catheter maintenance solution at 37 °C. The data obtained (Table 8.5) show that the incorporation of a surface active agent into the novel silicones increased the removal of both bacteria and encrusted deposits from the silicone surface when in contact with the benzalkonium chloride catheter patency solution compared with silicone elastomer without incorporated surfactant (A. D. Woolfson et al., unpublished data, 2008).

8.7

Toxicity and regulatory issues

8.7.1 Elastomer biocompatibility Lubricious RTV silicone elastomers should present no biocompatibility problems in terms of cytotoxicity, irritation or sensitisation. The silicone

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Table 8.5 Effect of 15 min contact with benzalkonium chloride (0.005% w/v) patency solution on surface encrustation and biofilm on the novel silicone elastomer with incorporated surface active agents Surface active agent incorporated

% surface active agent

% kill* of biofilm on material

Encrustation as surface cover (%)

Tween 20 Span 80 Lecithin None (control)

1.0 1.0 1.0 —

99.99 99.99 99.999 99.95

4.3 3.9 2.6 5.7

± ± ± ±

1.2 0.8 0.7 1.2

* Initial bacterial challenge in biofilm mode of growth: 5 × 105 colony forming units (cfu)/cm−2.

elastomer produced when PDMS is crosslinked with TODDOS is chemically identical to that produced in the reaction with TPOS. This material has a history of safe use in contact with epithelial tissue. It is presently used in the pharmaceutical industry for the manufacture of drug delivery devices. Thus, biocompatibility issues relate not to the elastomer but to the surface exudates that distinguish lubricious RTV silicone elastomers from their conventional counterparts. Several of the alternative lubricious alcohols produced in the crosslinking reaction, notably oleyl alcohol and octyldodecanol, are well known in the cosmetics and pharmaceutical industries. The safety of these oils has been detailed in a comprehensive published report (Anon., 1985).

8.7.2 Cytotoxicity testing Cytotoxicity or tissue culture testing was first included in the United States Pharmacopeia (USP) 22 in 1990. In 1993, the AAMI (Association for the Advancement of Medical Instrumentation) Biological Evaluation Committee adopted ISO 10993-5 as an American National Standard. The section, entitled ‘Biological evaluation of medical devices – Part 5: Tests for cytotoxicity: in vitro methods’, discusses methods used to evaluate the biological effect of medical devices using cells cultured under defined test conditions. In Part 1, ‘Guidance on selection of tests’, cytotoxicity testing is required for all device categories. The cell line typically used for cytotoxicity testing is the L929 mouse fibroblast cell. If appropriate for the end use of the product, other cell lines may be employed such as the rabbit cornea cells (SIRC) or embryonic lung cells (WI-38). There are three types of cytotoxicity tests: extract, direct contact and indirect contact. The nature of the sample and its intended use will determine which test is most appropriate. The extract test (minimum essential medium (MEM) elution test) is the most commonly used since it

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can be applied to a wide variety of raw materials and finished product configurations and extracts all exposed surfaces. The sample is extracted in the MEM used to culture the cells. The extract is then placed into contact with prepared cell cultures and the cells are observed microscopically after a defined incubation period. The cells are evaluated for signs of damage and reduced metabolic function and are graded according to a descriptive and numeric scale. In the MEM elution test, the commonly used sample extraction conditions are 37 °C for 24 hours; 48 hours after exposing the sample or extract to the cells, the monolayer is observed for cytotoxic effects. A numeric score is assigned based on the degree of change in cellular morphology and/or lysis. A score of zero indicates normal growth with little or no effect on the cell layer. A score of four represents nearly complete destruction of the cell layer. A score of two or less is considered to be a pass and a score of three or four indicates cytotoxicity. Positive and negative controls are always run to confirm the suitability of the test system. Cytotoxicy testing (MEM elution), carried out on conventional RTV silicone elastomer and ODD-derived silicone elastomer, produced zero scores in all cases compared with a maximum in the positive controls.

8.8

Conclusions

Silicone elastomeric devices display a high degree of biocompatibility and are quite flexible, yet mechanically strong. The material is often used for manufacturing urinary catheters. However, the lack of inherent lubricity of the material poses a major disadvantage as it causes discomfort and irritation to the patient. Self-lubricating silicone elastomers may be advantageous in this respect. Thus, novel tetra-functional compounds were developed for the condensation crosslinking of hydroxyl terminated PDMS chains to yield lubricious silicone elastomeric biomaterials. Lubricity was due to the higher alcohol, non-volatile condensation by-product formed during the crosslinking (vulcanisation) of linear PDMS. The alcohol, for example oleyl alcohol, was found to permeate to the surface and produce a liquid, oily film on the surface. The coefficients of friction of these elastomers were examined and it was found that the novel elastomers exhibited a greater lubricity, with substantially reduced coefficients of friction, when compared with a standard RTV silicone elastomer. Although the increase in lubricity may be of benefit when manufacturing a urinary catheter, the subsequent effect on the mechanical properties may result in the material not being suitable for its intended purpose. Therefore, the novel elastomers were also examined for their mechanical properties. From tensile analysis tests it was shown that, although their stiffness, as measured by the Young’s modulus, and ultimate strength at break were inferior when compared to the standard silicones, the ultimate strain at break was increased. This may be due to the condensation alcohol either

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interfering with the crosslinking reaction or its inherent volume decreasing the concentration of PDMS crosslinked per unit area. Both may result in a reduction in the overall crosslinking density. Nevertheless, good quality elastomers were produced. Biofilm formation on biomaterials in vivo is an additional, major factor that may adversely affect patients using urinary devices. Thus, the novel, lubricious silicones were further developed to demonstrate release of a model antibacterial drug, metronidazole. The lubricious silicone elastomers demonstrated enhanced release of this moderately hydrophilic agent compared with conventionally cross linked RTV silicone. Further surface treatment was made possible by incorporating a surfactant in the elastomer matrix, to aid removal of any adhering surface material, including, potentially, encrustation with inorganic salts in vivo. Finally, the lubricious silicones were shown to pass in vitro biocompatibility tests. Further work remains to be done, however, to demonstrate their practical utility and safety in the in vivo situation.

8.9

References

anon. (1985), ‘Toxicology report on oleyl alcohol and related higher alkanols’, J Am Coll Toxicol, 4, 1–29. arkles b (1983), ‘Look what you can make out of silicones’. Chem Tech, 13, 542–555. bradley d c, mehrotra r c and gaur d p (1978) Metal Alkoxides, London, Academic Press. braley s (1970), ‘The chemistry and properties of the medical grade silicones’, J Macromol Sci: Chem, A4, 529–544. chalk a j and harrod j f (1965), ‘Homogeneous catalysis II. The mechanism of the hydrosilation of olefins catalyzed by group Vii metal complexes’, J Am Chem Soc, 87, 16–21. colvin e w (1981), Silicon in Organic Chemistry, London, Butterworths. gorman s p and jones d s (2003), ‘Complications of urinary devices’, in: Wilson, M (Ed.), Medical Implications of Biofilms, Cambridge University Press, pp. 136–170. habal m b (1984), ‘The biologic basis for the clinical application of the silicones’, Silicone, 119, 843–848. jones d s, woolfson a d and djokic j (1996), ‘Texture profile analysis of bioadhesive polymeric semi-solids: mechanical characterisation and investigation of interactions betwen formulation components’, J Appl Polymer Sci, 61, 2229–2234. jones d s, garvin c p and gorman s p (2004), ‘Relationship between biomedical catheter surface properties and lubricity as determined using textural analysis and multiple regression analysis’, Biomaterials, 25, 1421–1428. khoury a e, olson m e, villari f and costerton j w (1991), ‘Determination of the coefficient of kinetic friction of urinary catheter materials’, J Urol, 145, 610–612.

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kunin c m (1989), ‘Blockage of urinary catheters: role of microorganisms and constituents of the urine on formation of encrustation’, J Clin Epidemiol, 42, 835–842. levin r (1958), The Pharmacy of Silicones and their Uses in Medicine, London, Morgan Brothers. malcolm r k, mccullagh s d, woolfson a d, gorman s p, jones d s and cuddy j (2004), ‘Controlled release of a model antibacterial drugfrom a novel selflubricating silicone biomaterial’, J Control Release, 97, 313–320. mccullagh s d, malcolm r k, woolfson a d, gorman s p, jones d s and cuddy j (2004), ‘Release kinetics of oleyl alcohol from a self-lubricating silicone biomaterial’, J Mater Chem, 14, 1093–1098. nicolle l e (2005), ‘Catheter-related urinary tract infection’, Drugs Aging, 22, 627–639. polmanteer k e (1981), ‘Current perspectives on silicone rubber technology’, Rubber Chem Technol, 54, 1051–1080. ratner b d (2004), Biomaterials Sciences: An Introduction to Materials in Medicine, 2nd edn, New York, Academic Press. ritger p l and peppas n a (1987), ‘A simple equation for description of solute release: I. Fickian and non-Fickian release from non swellable devices in the form of slabs, spheres, cylinders or discs’, J Control Release, 5, 23–36. speier j l, webster j a and barnes g (1957), ‘Addition of silicone hydrides to olefinic double bonds: Part II: Use of Group VIII metal catalysts’, J Am Chem Soc, 79, 974–981. stickler d j, morris n, moreno n c and sabbuba n (1998), ‘Studies on the formation of crystalline bacterial biofilms on urethral catheters’, Eur J Clin Microbiol Infect Dis, 18, 206–208. tunney m m and gorman s p (2002), ‘Evaluation of a poly(vinyl pyrollidone)-coated biomaterial for urological use’, Biomaterials, 23, 4601–4608. van der weijj f w (1980), ‘The action of tin compounds in condensation-type RTV silicone rubbers’, Makromolecule Chemie, 181, 2541–2548. woolfson a d, malcolm r k, gorman s p, jones d s, brown a f and mccullagh s d (2003), ‘Self-lubricating silicone elastomer biomaterials’, J Mater Chem, 13, 2465–2470.

9 Temporary urethral stents T. TA M M E L A, Tampere University Hospital, Finland

Abstract: Temporary urethral stents are designed for short-term use, usually for between 4 months and 3 years. Indications for use include treatment of urethral strictures, prostatic obstruction and detrusorsphincter dyssynergia. An ideal temporary stent should enable easy placement under local anaesthesia, have minimal local side-effects (such as tissue hyperplasia and encrustation) and low risk of migration; it should also be easy to remove. The development of both metallic and polymer stents has resulted in better functioning devices that can be removed more easily but this still necessitates some kind of procedure. In addition, encrustation and migration remain as significant problems. An ideal biodegradable stent would provide adequate support to the duct wall, keep the lumen open during and after the healing process and biodegrade completely from the body. Polylactide and polyglycolide are the most commonly used biodegradable materials. Good mechanical properties can be achieved by self-reinforcing procedures where molecular chains form the polymer matrix. The stents can be made self-expanding and their degradation time can vary between 2 weeks and 1 year. However, there is a need to develop new stent designs that have preferable bioactivity and thereby are able to modulate the healing process. Key words: spirals, biodegradation, retrievable stents, migration, encrustation, prostatic obstruction, urethral stricture.

9.1

Introduction

Since 1980 different types of temporary and permanent stents have been introduced into urological practice to release obstruction of the prostatic urethra or the external sphincter and for the treatment of recurrent urethral strictures after internal urethrotomy (Fabian 1980). The devices are categorized as permanent or temporary, with the temporary stents being designed primarily for short-term use, usually for between 4 months and 3 years (Yachia 1997, Shin et al. 2006). The characteristics of an ideal temporary stent include easy placement under local anaesthesia, minimal local sideeffects – such as tissue hyperplasia or encrustation – and a low risk of migration. In addition, endoscopy through the stent should be feasible and the stent must be easily removable or, preferably, biodegradable so that no 208

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further procedure in necessary. This chapter will dicuss both retrievable biostable stents and biodegradable stents, with a special emphasis on the latter, including indications and complications associated with their use.

9.2

Indications for the use of stents

9.2.1 Urethral stricture Urethral stricture is one of the oldest known urological diseases and remains a common problem associated with high morbidity (Yelderman and Weaver 1967, Steenkamp et al. 1997). The natural history of the condition usually begins with a lesion of the urethral mucosa and infection followed by a scar. The main types of urethral strictures are iatrogenic, inflammatory and traumatic together with those of unknown aetiology. Dilatation and direct-vision internal urethrotomy have been the standard treatments while different kinds of urethroplasty techniques have been used in the most difficult cases. After endoscopic treatment, success rates of 50–60% can be obtained in short strictures without spongiofibrosis while in longer or multiple strictures involving the corpus spongiosum the recurrence rate is much higher. Higher success rates are also achieved with an iatrogenic stricture than with those of post-traumatic and postinflammatory aetiology (Stone et al. 1983). Generally it is advisable to avoid more than two internal urethrotomies in the primary treatment of urethral strictures as the procedure induces a local inflammatory reaction which frequently develops a more extensive stricture than the treated one (Albers et al. 1996). Cure rates have not been improved by repeated multiple incisions in the circumference of the stricture, glucocorticoid injection into the scar, use of indwelling catheters or postoperative hydraulic self-dilatation. The essential problem is how to prevent the edges of the cut stricture from adhering together and the scar from shrinking. Many patients do not accept the option of repeated self-catheterization to maintain satisfactory urethral lumen after internal urethrotomy (Niesel et al. 1995). In order to solve the problem of recurrence, permanent self-expanding metallic stents, temporarily placed biocompatible metallic stents and polyurethane temporary stents have been introduced for the treatment of urethral strictures. However, problems with epithelial hypertrophic proliferation, and stent encrustation and difficulties with removal have been associated with the permanent and temporary endoscopically placed urethral stents used so far (Kletscher and Oesterling 1994). There are a wide range of surgical techniques for the treatment of urethral stricture including excision of the stricture and replacing or patching the stenosed segment of the urethra with tissues taken from other parts of

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the body such as skin, bladder and buccal mucosa. Although encouraging results have been reported these techniques only work in experienced hands.

9.2.2 Prostatic obstruction During the last 15 years a number of mini-invasive treatments for benign prostatic enlargement have been introduced, including interstitial laser coagulation (ILC), visual laser ablation of the prostate (VLAP) and highenergy transurethral microwave thermotherapy (TUMT). Most of these treatments induce tissue oedema in the prostate which increases bladder outlet obstruction (BOO) and results in postoperative urinary retention requiring permanent or intermittent catheterization for up to 6–8 weeks (Madersbacher 2005). Similarly 10–15% of patients undergoing brachytherapy develop urinary retention either immediately or in a few days following implantation (Terk et al. 1998). For diagnostic purposes stents can be used to predict the outcome of transurethral resection of the prostate (TURP) in difficult cases, for example cases where there is a combination of benign prostatic obstruction, severe detrusor overactivity and urgency urinary incontinence. In patients suffering from Parkinson’s disease or multiple sclerosis, the risk of de novo or exacerbated postoperative urgency incontinence is substantial. Temporary stents could also be used following acute urinary retention to keep the prostatic urethra open while the size of the prostate is reduced by using 5-alpha-reductase inhibitors such as finasteride or dutasteride. Similarly, they can be used while waiting for TURP in cases where comorbidities necessitate postponment of surgery or where there is a long waiting list.

9.2.3 Detrusor-sphincter dyssynergia Traumatic, inflammatory, neoplastic, vascular or congenital suprasacral spinal cord lesions are responsible for voiding disorders due to detrusor overactivity and detrusor-sphincter dyssynergia (DSD). The functional obstruction induced by these disorders results in a high-pressure system that can be responsible for genitourinary tract complications. The primary and effective way to treat the condition is drainage with intermittent selfcatheterization. An alternative, particularly for those with poor dexterity, is to reduce the bladder outlet resistance. Transurethral sphincterotomy has been the standard operation to reduce the outflow resistance but it is associated with complications such as profuse bleeding and is not always successful. Temporary stents could therefore be used to evaluate whether the patient would gain any benefit from reduction of the outlet obstruction.

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It would also be beneficial in cases when the final outcome of the suprasacral disease causing the lower urinary tract disorder is not clear.

9.3

Non-degradable temporary urethral stents

Available biostable stents are made of various materials, including stainless steel, nickel–titanium alloy (nitinol) and polymers. The temporary stents should either be removed or changed every 4 to 36 months (Yachia 1997, Shin et al. 2006).

9.3.1 Metallic stents First-generation stents Metallic temporary stents are made of coiled wire. The first-generation spiral stents were made of stainless steel and were based on the original Fabian stent. They were developed for prostatic obstructions and examples include the stainless steel Urospiral®, the gold-plated Prostakath® and the nickel–titanium ProstaCoil®. All three have a long segment to hold the prostatic lumen open, a bulbar segment to anchor the stent in place and a transsphincteric spacer connecting the prostatic and bulbar segments (Yachia 1997). Second-generation stents The second-generation spiral stents are made of an alloy of nickel and titanium and are either self- or thermo-expandable (ProstaCoil®, Memokath®). Memokath 028® is a single-segment device and upon insertion the entire outer calibre of the stent is 22 F, but when it is flushed with 50 °C water the stent expands to its memorized shape and anchors itself in place (Ellis and Gidlow 1996). The Memokath 028® is a tightly coiled stent that is designed to prevent urothelial ingrowth. When cooled using water at 10 °C, it becomes soft and pliable, and can be removed with relative ease under local urethral anaesthesia. The alloy is not magnetic and will therefore not preclude the patient from magnetic resonance imaging (MRI). It also has excellent biocompatibility properties. Yachia and Beyar (1991, 1993) introduced a temporary but long-term stent for the treatment of urethral strictures. The new concept involved placement of a stent in the stenotic urethra after dilating or incising the stricture, and leaving the stent in place long enough to keep the lumen open until stabilization of the scarring process, after which it could be removed. The presence of the stent in the lumen aimed to prevent scar contraction during the healing process. The temporary, large-calibre

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UroCoil® system stents were made of nickel–titanium alloy. They had an expanded calibre of 24–30 F with an insertion calibre of 17 F. In the UroCoil® system different stents were developed for use in the different parts of the urethra, including a stent for penile urethra strictures, a stent for bulbar urethral strictures and a stent for combined prostatic and bulbar urethral strictures. They have been left indwelling for 12 months and reported recurrence rates were 17, 20 and 20% at 2, 3 and 4 years respectively. Later the thermo-expandable Memokath® urethral stent made of nitinol was launched, also intended for the treatment of recurrent urethral strictures (Yachia 1997). Covered metallic stents Covered retrievable expandable nitinol stents have recently been introduced for treatment of recurrent urethral strictures (Song et al. 2003, Shin et al. 2006). The Song urethral stent is woven from a single thread of 0.1-mm-diameter nitinol filament in a tubular configuration and is covered in polyurethane or polytetrafluoroethylene to prevent tissue hyperplasia ingrowth through the stent wires into the lumen. In order to enable removal of the stent, 2-mm-diameter nylon loops are hooked inside each bend of the distal end of the stent and nylon threads are passed through each of the nylon loops (drawstrings); these threads fill the inner circumference of the distal stent. Clinical outcome Temporary metallic stents are mostly inserted under direct cystoscopic vision using urethral local anaesthesia. The insertion of the Memokath® stent is no more uncomfortable than cystoscopy until the instillation of hot water into the urethra and bladder, which causes discomfort for 4–5 s. Most men who had the Memokath® stent inserted for bladder outlet obstruction to the prostatic urethra voided immediately after the procedure so that the outcome of the stent placement was immediately apparent (Perry et al. 2002). Although the target voiding parameters improved immediately after insertion of a nitinol spiral stent, the maximum improvement in quality of life was not achieved until 1 month after placement. This might be because the stent caused irritation in the trigone (van Dijk et al. 2005). When compared with long-term indwelling catheterization, urinary infection was significantly less frequent and less severe in patients with urethral stents (Egilmez et al. 2006). It was concluded that the Memokath® intraprostatic stent was a valuable addition to the armamentarium of the urologist treating elderly or frail men with advanced bladder outlet obstruction and that it complements existing technologies (Perry et al. 2002). However, the

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permanent use of stents to relieve benign prostatic obstruction is not indicated in elective settings (Madersbacher et al. 2004). Placement of a covered retrievable stent for a minimum of 4 months was effective in inducing long-term resolution of refractory urethral strictures. All patients voided well after stent placement. All reported mild urgency and discomfort at the site of stent placement; these problems resolved spontaneously within 1 week after stent placement (Choi et al. 2007). In patients suffering from DSD, temporary urethral sphincter stents ensured effective bladder emptying and prevented autonomous and infectious complications (Game et al. 2008). After temporary stenting, most patients were subsequently treated with permanent urethral sphincter stents, while several patients were able to start intermittent selfcatheterization later. Complications associated with temporary metallic stents One of the problems with temporary self-expanding metallic stents is the shortening that occurs; this may be 10–50% of the stent length depending on the type of stent used. Other common problems are migration of the stent and tissue proliferation into the stent lumen. Migration was particularly a problem in the non-expandable, first-generation spiral stents (10–38% of cases), but reports on other temporary stents also showed varying high migration rates. The thermoexpandable spiral stents have been reported to migrate less often (0–13%) (van Dijk et al. 2005). However, even in the newest covered retrievable expandable nitinol stents, migration remains the largest obstacle to achieving the adequate duration of stent placement critical for achieving long-term resolution (Choi et al. 2007). In addition, haematuria is a relatively common complication. The haematuria directly after the insertion procedure is the result of small lesions of the urethra caused by manipulation during delivery. An explanation for the transient haematuria occurring later is probably that physical activity causes friction of the stent within the prostatic urethra, leading to damage of the urethral lining. In particular, stents used for treatment of urethral stricture may develop sphincteric dysfunction when they are deployed in the near vicinity of the external sphincter, and they may also induce tissue proliferation at the ends of the stents. When the stent is located in the anterior urethra it often causes postmicturition dribbling. Tissue hyperplasia of the urothelium results in luminal narrowing which may cause obstructive symptoms and may necessitate premature removal of the stent. Another relatively common reason for the development of obstruction is encrustation with calculus formation, which seems to be more of a problem in patients with spinal cord injuries (Perry et al. 2002, Mehta and Tophill 2006). Infected urine and high residual urine volumes

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predispose to encrustation and calculus formation. Obviously there are also differences in the characteristics of the different materials used in the stents. From a clinical point of view, encrustation is a significant problem as it makes removal of the retrievable stents more difficult and causes damage to the urethra.

9.3.2 Polymer stents As an alternative to metallic stents, several polyurethane stents have been developed. In 1995, Nissenkorn (1995) introduced a polyurethane intraurethral catheter that comprised a tubular device with basket dilatation at both ends of the stent, at the bladder neck and at the apex of the prostate. Nissenkorn and Shalev (1997) also treated urethral stricture patients with a temporary polyurethane stent. They recommended replacement of the stent once a year in order to prevent encrustations and obstruction. The Barnes stent is also a polyurethane device in which the proximal end is similar in design to a urethral catheter, whereas distally a single retaining basket is designed to sit at the verumontanum (Barnes and Yakubu 1998). The Trestle catheter consists of two silicone tubes with a thread connection. The proximal prostatic part has the Foley catheter design and the distal tube is in the bulbous urethra. The Trestle intraurethral catheter has been used in temporary stenting of urethral after high-energy transurethral microwave therapy of the prostate (Devonec and Dahlstrand 1998). Recently, a novel polyurethane stent (Spanner®) was presented. This is also similar to the proximal 4–6 cm of a Foley catheter, including a proximal balloon to prevent distal placement, a urine port situated cephalad to the balloon and a reinforced stent of various lengths to span the prostatic urethra (Corica et al. 2004). It is easily inserted and removed, remains anchored in position and significantly improves voiding in patients who have obstruction in the prostatic urethra.

9.4

Biodegradable urethral stents

9.4.1 Requirements for a biodegradable stent In 1993, the first biodegradable self-reinforced poly-l-lactic acid (SRPLLA) urethral stent was introduced; following this, the development of new materials and configurations of urological stents has been rapid (Kemppainen et al. 1993). An ideal device in the lower urinary tract would provide adequate support to the duct wall, like the urethra, keep the lumen open during and after the healing process and biodegrade completely from the body. The material must fulfil certain biocompatibility demands according to the guidelines for the tissue biocombatibility analysis

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and risk assessment of new medical devices. The rigidity of the material must be suitable for operating targets, the degradation products must also be biocompatible metabolic products and the rate of degradation suitable for tissue regeneration. The devices must also have good sterilization properties.

9.4.2 Bioabsorption and biodegradation The term ‘bioabsorbtion’ means the degradation and metabolism of the material in vivo to small molecules, like carbon dioxide and water, and to energy. The term ‘biodegradation’ means the morphological and chemical degradation of the material in vivo, and the term ‘degradation’ represents the general breakdown of the material. The stents used for treatment of urethral strictures in the anterior urethra mostly bioabsorb by being implanted into the tissue, whereas those used in the prostatic urethra biodegrade into small fragments which are then washed out with urine. Degradation of the bioabsorbable polymers is the sum of many factors and is accelerated by the presence of residual monomers and oligomers in the polymer, the alkalinic pH of the surroundings, reduction of crystallinity and orientation, certain enzymes – e.g. pronase, proteinase K and bromeline – as well as the sites of implantation where active metabolism, many muscular movements and stresses are loading the implant.

9.4.3 Biodegradable materials Polylactide (PLA, polylactic acid) and polyglycolide (PGA, polyglycolic acid) are the most commonly used biodegradable materials. They belong to the group of poly-alpha-hydroxy acids that are members of the polyester family. Lactic acid (2-hydroxypropanoic acid) CH3CHOHCOOH has two enantiomers, l(+)-lactic acid and d(−)-lactic acid, which differ significantly in their rates of biodegradation (Cutright et al. 1974). The degradation rate increases when the proportion of d-lactide increases from 0 to 50 mol % (Kulkarni et al. 1966). For example, PLA 96/4 is a polymer of l- and dlactide with a ratio of 96/4 l- and d-lactic acid, respectively. In addition, the physical properties of the copolymers depend on the relative amounts of the l and d configurations (Vert et al. 1992). Glycolic acid (hydroxyethanoic acid) HOCH2COOH is synthesized by ring opening polymerization from glycolide, resulting in a poly-alphahydroxy derivate. Glycolic acid has no chiral centre in the molecule and therefore it does not form enantiomers (Gilding and Reed 1979). Polyglycolide was the first commercially successful synthetic biodegradable polymer to be used as a biomedical material. The great interest in using PGA as a device material is due to its good mechanical and absorbtion

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properties. The degradation products are metabolic products that normally occur in living tissue, e.g. polylactides. The reaction that produces a polymer from monomers is known as polymerization. A combination of two different monomers can produce a random polymer, a block copolymer or a graft copolymer. In 1975, Vicryl became the first commercially available copolymer of PLA and polyglycolic acid and was used for surgical sutures (Vainonpää et al. 1989). PLGA is a copolymer of lactic and glycolic acid, mostly used with an 80/20 ratio of PLA/PGA. Copolymers are less crystalline than their constituent homopolymers and consequently degrade more rapidly; an increase in the amount of PGA in PLGA accelerates degradation.

9.4.4 Mechanical properties and degradation of self-reinforced composite devices PGA and PLA polymer devices are partially crystalline, linear-chain degradable polymers and show only modest values for mechanical strength (Vainonpää et al. 1989). Good mechanical properties can be achieved by using extrusion and die-drawing techniques (Törmälä 1992). During this self-reinforcing procedure, aligned molecular chains are formed in the material from the polymer matrix. These aligned chains form fibrils and filaments, which carry the forces directed to the implant. Therefore, the fibrous material and the matrix material have the same chemical element composition. When the molecular microstructure is oriented, the mechanical strength, modulus and toughness of the absorbable polymers increase significantly (Törmälä 1992). The mechanical properties are also dependent on the basic molecule and the length of the polymer chains. In addition, the number of mono- and oligomer residues of polymerization, the configuration and the total mass of the material are important factors in the degradation process (Välimaa et al. 1998). SR biodegradable stents can be made self-expanding at body temperature as a result of the viscoelastic memory of the material. The ability to decide the speed and rate of expansion of stents is of great clinical importance because of the varying requirements of the different indications for use of stents. The level at which the expansion stops, as well as the speed of the expansion, depend on the material, crystallinity, internal arrangement of molecular chains, initial diameter of the spiral, diameter of the stent wire and annealing temperature (Välimaa et al. 2002). The time required for expansion of the stent to its final diameter may range from 5 minutes to 2 weeks (Välimaa et al. 1998, Välimaa et al. 2002). In most urological indications rapid expansion would be preferable. By using a new tubular mesh

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9.1 Braided PLA biodegradable urethral stent (outer diameter 8 mm).

configuration the rapid expansion properties can be further improved (Fig. 9.1) (Vaajanen et al. 2003). The strength retention time and the functional time of the biodegradable implant varies according to the material selected, the molecular weight, the molecular weight distribution and the morphology of the polymer, and also the distribution of repeat units. In addition, the process parameters and the sterilization method also have their own effects on the molecular weight, molecular weight distribution and morphology of the polymer (Törmälä et al. 1998). The degradation rate of the material depends on the moisture content and the temperature; therefore, the processing, sterilization and storage of these materials should take place in a dry atmosphere and at as low a temperature as possible (Välimaa et al. 2002). The biodegradable SR stents have been sterilized either by ethylene oxide or by gamma irradiation methods. The water solubility of lactide and glycolide monomer is very high. By hydrolysis, PLA and PGA degrade first into short molecular chains (oligomers) and then into basic acids. PLLA degrades into l-lactic acid, poly-d-lactic acid (PDLA) into d-lactic acid and PGA into glycolic acid (McNeill and Leiper 1985). The degradation times of the different copolymers are shown in Table 9.1.

9.4.5 Biocompatibility Bioabsorbable polymers PLA and PGA have been proven to have good biocompatibility properties as suture materials over a period of 30 years. PGA is well tolerated by the soft tissue, evoking only minimal inflammatory response (Herrmann et al. 1970). The biodegradation of macroscopical SR-PGA implants proceeds by a cellular reaction comparable with the biodegradation reactions of polyglycolide sutures (Echeverria and Jimenez 1970). SR-PLLA spirals have shown good biocompatibility with minimal tissue reaction around the stent in the anterior urethra (Kemppainen et al. 1993). Biodegradable materials were also found to have similar biocombatibility to that of silicone, and better biocompatibility than latex in an animal toxicity test (Laaksovirta et al. 2002c).

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Table 9.1 Degradation times of different bioabsorbable copolymers in vitro and in vivo

Name of the copolymer

Abbreviation

Degradation time Degradation in vitro time in vivo

Polyglycolic acid Poly-L-lactic acid Poly-L,D-lactic acid Polylactic–glycolic acid

PGA PLLA PLA 96L/4D PLGA 80L/20G

3–4 weeks 12 months 30 weeks 2 months

3–4 weeks 8–10 months 5–6 months 2–3 months

9.4.6 Bacterial adherence PGA or PLA materials have no antibacterial properties and the extent of bacterial adherence has been shown to correlate more with the bacterial strain than with the material of the stent (Cormio et al. 1997). It has been suggested that dissolution of the underlying material which the bacteria have been attached to during degradation, results in release of bacteria. However, both silver nitrate and ofloxacine blended poly (ε-caprolactone) homopolymer coatings prevented the adherence of bacteria to PLA stent pieces in vitro. The prevention effect correlated with the concentration of antibacterial agent in the homopolymer coating (Petas et al. 1998, Multanen et al. 2000).

9.4.7 Encrustation SR-PLLA stents had fewer encrustations than stainless steel stents in the anterior urethra of the rabbit (Kemppainen et al. 1993). Likewise, there was some encrustation on gold-plated steel wire (Prostakath®), but none on SR-PGA or SR-PLA96/4 stents after incubation for 2 weeks in artificial urine (Cormio et al. 1997, Petas et al. 1997b). In addition, a SR-PLGA80/20 stent was shown to be markedly more resistant to encrustation than metallic urethral stents (Laaksovirta et al. 2002b). The absence of encrustation in the biodegradable material can be explained by the sloughing off of the surface of a biodegradable stent as a result of the continuous hydrolysation.

9.4.8 Biodegradable stents Biodegradable prostatic stents Self-expandable spiral prostatic stents have been developed from different biodegradable materials, including SR-PGA, SR-PLLA, SR-PLA96/4 with

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barium and SR-PLGA. The configuration of the spiral stent resembles that of the Fabian spiral stent (Fabian 1980) (Fig. 9.1). The outer diameter of the spiral is 7 or 8 mm (Ch 21 or 24), and the prostatic portion is 45–65 mm long. The neck of the spiral is 20 mm in length. The biodegradable SR-PGA spiral stents have been used with favourable results after VLAP (Petas et al. 1997c), ILC of the prostate (Petas et al. 2000) and TUMT (Dahlstrand et al. 1997). At body temperature, the outer diameter of the spiral stent increases by more than 60% which fixes it in place. The spiral is inserted by pushing it into the prostatic urethra with the tip of the cystoscope immediately after the laser therapy and the patients are allowed to void immediately. When a SR-PGA stent was inserted, spontaneous voiding started in approximately 90% of cases compared with 35% in the group having only a suprapubic catheter after VLAP (Petas et al. 1997c, Petas et al. 2000). However, the strength retention time of the SR-PGA stents was too short for some patients who reported diminished force of the urinary flow stream at the time of degradation at 3–4 weeks. The total urinary infection rate was lower in the spiral stent group (20% of patients) than in the suprapubic (35%) or indwelling catheter groups (36%). The Danish Prostatic Symptom Score 1 (DAN-PSS1) weighted symptoms and peak flow rates of the SR-PGA spiral stent groups were comparable with those of the two other groups (Petas et al. 1997c, Petas et al. 2000). SR-PGA spiral stents have also been used successfully in combination with high-energy TUMT therapy to prevent postoperative urinary retention (Dahlstrand et al. 1997) as well as being used as a tool to test the risk for post-TURP incontinence in patients with combined benign prostatic obstruction and severe bladder overactivity (Knutson et al. 2002). The in vitro degradation of SR-PLA 96/4 spiral stent takes 30 weeks (Välimaa and Törmälä 1996) and it has also been shown to prevent urinary retention effectively in patients undergoing VLAP due to benign prostatic obstruction. The mean degradation time of the SR-PLA 96/4 stent in clinical use was 6 months (Petas et al. 1997a). With the aim of developing a biodegradable prostatic stent that would have a degradation time of approximately 2 months, an SR-PLGA spiral stent was constructed. In in vitro experiments the expansion started rapidly during the first 4 hours and then continued more slowly over the next 3 days up to 100% (Välimaa et al. 2002). In clinical studies where patients underwent ILC of the prostate for benign prostate enlargement (BPE) and an SR-PLGA 80/20 spiral stent was inserted upon completion of the operation, more than 90% of patients were able to void on the first postoperative day (Laaksovirta et al. 2002a). In two cases only the stent had moved proximally and had to be relocated, whereafter voiding was successful. At 2 months the stent was still present intact in the urethra in all except three patients,

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but at 4 months it had degraded into small fragments and at 6 months had been completely eliminated. Some 50% of the patients had irritative symptoms caused at least partly by the ILC of the prostate itself; 10% had asymptomatic urinary tract infection postoperatively. In all patients the degradation time was long enough to meet the requirements for postprocedure urinary drainage. Biodegradable stents have also been tested in combination with prostate size-decreasing finasteride therapy in patients with acute urinary retention (AUR) owing to BPE. SR-PLLA was chosen as the stent material because it has approximately a 1-year biodegradation time, which leaves sufficient time for finasteride therapy to achieve its maximum effect. In an open, nonrandomized pilot study the combination treatment in 11 men with AUR as a result of BPE gave promising results; all patients started to void spontaneously within 2 weeks and during the mean follow-up of 24 months, only 3 patients required surgical treatment (Isotalo et al. 2000). However, in a randomized placebo-controlled study enrolling 55 patients with AUR due to BPE, where the indication for the combination therapy was an impaired general condition that would have increased risk in surgery, only 19 patients completed the study (Isotalo et al. 2001). The main reason for discontinuation within the first 6 months was insufficient therapeutic response, whereas after this period the rapid uncontrolled breakdown of the stent configuration caused obstruction resulting in impaired voiding and increase of lower urinary tract symptoms. Biodegradable urethral stents for urethral strictures Despite the development of endoscopic and reconstructive urology and the use of different urethral stents, the recurrence rate of urethral strictures remains a problem. The essential problem is how to prevent the edges of the cut stricture from adhering together and the scar from shrinking after urethrotomy. A biodegradable SR-PLLA stent with helical spiral configuration was developed to avoid these problems (Isotalo et al. 1998). PLLA was chosen as the device material because of its degradation properties and its long 1-year degradation time. In the early experiments the spiral stent had no expansion properties and it had to be fixed with sutures into the urethral wall; therefore its clinical usefulness was reduced. As a result of developments in manufacturing technology an expandable spiral stent was introduced. The outer diameter of the stent increases 70% from the initial 8 mm (Ch 24). The expansion is fast and most of it is achieved during the first 30 minutes; there is therefore no need for external fixation of the stent. Although the SR-PLLA spirals were totally covered by urethral epithelium after 6 months, its clinical use was, however, limited by the risk of a sudden breakdown of the helical spiral configuration of the stent, leading to

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transient obstruction of the lower urinary tract (Isotalo et al. 1998, Isotalo et al. 2002). In order to overcome these problems a new tubular mesh configuration was developed (Fig. 9.1). This further improves the expansion and degradation properties of the biodegradable urethral stent (Vaajanen et al. 2003, Isotalo et al. 2005, Isotalo et al. 2006). The mesh stent, however, still needs to be evaluated both experimentally and clinically. Biodegradable SR-PGA spiral stents have also been used in free skin urethroplasty for bulbous urethral strictures. The stricture is opened by internal urethrotomy and the preputial free skin graft is mounted over the spiral stent. The graft is located endoscopically and fixed with percutaneous sutures at the site of the strictured urethra. Suprapubic urinary diversion is used for 10 days. The preliminary results were successful in patients with recurrent bulbous urethral strictures. The need for a long hospitalization period has, however, reduced the acceptability of this method (Oesterlink and Talja 2000).

9.5

Future trends

Although the results of using either biostable or biodegradable stents in the treatment of recurrent strictures, especially in the bulbous urethra, are in some respects encouraging, the number of failures is still high. An important reason for this is apparently the excessive urethral scarring and periurethral fibrosis in the patients with the most chronic recurrent urethral strictures. Because a stent cannot prevent recurrence of urethral stricture after urethrotomy, at least in the most difficult cases, it is necessary to develop bioactive retrievable and biodegradable stents that could modulate formation of the scar tissue. Although the biodegradable stent offers a new and, for the patient, convenient option for avoiding the use of an indwelling catheter in procedures that cause oedema and postoperative urinary retention, the price of the stent may be a limiting factor in the acceptance of the treatment modality. Irritative symptoms are common to all stents and cannot be completely avoided when biodegradable stents are used. In order to minimize the irritating effects of a degrading stent, as well as the effects of its sudden collapse in the terminal phase of biodegradation, new configurations of biodegradable urethral stents should be developed and compared with the retrievable stents available or being developed. The newest design of biodegradable urethral stents is a tubular helical mesh. The preliminary experimental experiences have been encouraging, but clinical trials are still needed. Similarly there is a need for smoother and softer retrievable stents. At the moment it appears that covered retrievable stents will become increasingly popular as they remain in the urethral lumen without becoming epithelialized.

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Further experimental as well as controlled clinical studies are needed to investigate the new design stents with preferable bioactivity; these stents should be compared with other methods for preventing urinary retention after the types of BPE therapies and brachytherapy that induce prostatic oedema as well as with methods for the temporary treatment of patients with urinary retention awaiting surgery. Similarly, controlled studies are needed to compare the use of biodegradable stents in the treatment of urethral strictures with other forms of therapy. An intensive study to discover the best possible materials, models, coating materials and additives for both biodegradable and retrievable stents is continuing. The use of biomaterials will also bring new therapy methods for future clinical use.

9.6

References

albers p, fichtner j, bruhl p and muller sc (1996), Long-term results of internal urethrotomy, J Urol, 156, 1611–14. barnes dg and yakubu a (1998), Temporary prostatic stenting using the Barnes stent. In: Stenting the Urinary Tract, ed. Yachia D, Isis Medical Media, Oxford, pp. 335–8. choi ek, song hy, shin jh, lim jo, park h and kim cs (2007), Management of recurrent urethral strictures with covered retrievable expandable nitinol stents: longterm results, AJR, 188, 1517–22. corica ap, larson bt, sagaz a, corica ag and larson tr (2004), A novel temporary stent for relief of prostatic urethral obstruction, BJU Int, 93, 346–8. cormio l, laforgia b, sitonen m, ruutu m, törmälä p and talma m (1997), Immersion in antibiotic solution prevents bacterial adhesion onto biodegradable stents, Br J Urol, 79, 409–13. cutright de, perez b, beasley jd, larson wj and posey wr (1974), Degradation rates of polymers and copolymers of polylactic and plyglycolic acids, Oral Surg, 37, 142–52. dahlstrand c, grundman s and petterson s (1997), High-energy transurethral microwave therapy for large severely obstructing prostates and the use of biodegradable stents to avoid catheterization after treatment, Br J Urol, 79, 907–9. devonec m and dahlstrand c (1998), Temporary urethral stenting after highenergy transurethral microwave thermotherapy of the prostate, World J Urol, 16, 120–3. echeverria e and jimenez j (1970), Evaluation of an absorbable synthetic suture material, Surg Gyn & Obs, 131, 1–14. egilmez t, aridogan a, yachia d and hassin d (2006), Comparison of nitinol urethral stent infections with indwelling catheter-associated urinary-tract infections, J Endourol, 20, 272–7. ellis bw and gidlow ab (1996), Thermoexpandable prostatic stents in frail or elderly men, a risk free and worthwhile technique, Eur Urol, 30 (suppl. 2), 110. fabian km (1980), Der intraprostatische ‘partielle Katheter’ (urologische Spirale), Urologe A, 19, 236.

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game x, chartier-kastler e, ayoub n, even-schneider a, richard f and denys p (2008), Outcome after treatment of detrusor-sphincter dyssynergia by temporary stent, Spinal Cord, 46, 74–7. gilding dk and reed am (1979), Biodegradable polymers for use in surgerypoly(glycolic)-poly(lactic acid) homo and copolymers, Polymer, 20, 1459–64. herrmann jb, kelly rj, and higgins ga (1970), Polyglycolic acid sutures, Arch Surg, 100, 486–90. isotalo t, tammela tlj, talja m, välimaa t and törmälä p (1998), A bioabsorbable self-expandable, self-reinforced poly-L-lactic acid urethral stent for recurrent urethral strictures: A preliminary report, J Urol, 160, 2033–6. isotalo t, talja m, välimaa t, törmälä p and tammela tlj (2000), A pilot study of a bioabsorbable self-reinforced poly L-lactic acid urethral stent combined with finasteride in the treatment of acute urinary retention from benign prostatic enlargement, BJU Int, 85, 83–6. isotalo, talja m, hellström p, perttilä i, välimaa t, törmälä p and tammela tlj (2001), A double-blind, randomized, placebo-controlled pilot study to investigate the effects of finasteride combined with a biodegradable selfreinforced poly L-lactide acid spiral stent in patients with urinary retention caused by bladder outlet obstruction from benign prostatic hyperplasia, BJU Int, 88, 30–4. isotalo t, talja m, törmälä p and tammela tlj (2002), A biobsorbable selfexpandable, self-reinforced poly-L-lactic acid urethral stent for recurrent urethral strictures; long term results, J Endourol, 16, 759–62. isotalo t, nuutinen j-p, vaajanen a, martikainen pm, laurila m, törmälä p, talja m and tammela tlj (2005), Biocompatibility and impantation properties of 2 differently braided, biodegradable, self-reinforced polylactic acid urethral stents: An experimental study in the rabbit, J Urol, 174, 2401–4. isotalo tm, nuutinen jp, vaajanen a, martikainen pm, laurila m, törmälä p, talja m and tammela tl (2006), Biocompatibility properties of a new braided biodegradable urethral stent: a comparison with a biodegradable spiral and a braided metallic stent in the rabbit urethra, BJU Int, 97, 856–9. kemppainen e, talja m, riihelä m, pohjonen t, törmälä p and alfthan o (1993), A bioresorbable urethral stent, An experimental study, Urol Res, 21, 235–8. kletscher ba and oesterling je (1994), Urethral stents: current status for the treatment of recurrent bulbar urethral strictures and benign prostatic hyperplasia, Curr Opin Urol, 4, 162–7. knutson t, pettersson s and dahlstrand c (2002), The use of biodegradable PGA stents to judge the risk of post-TURP incontinence in patients with combined bladder outlet obstruction and overactive bladder, Eur Urol, 42, 262–7. kulkarni rk, pani kc, neuman c and leonard f (1966), Polylactic acid for surgical implant, Arch Surg, 93, 839–43. laaksovirta s, isotalo t, talja m, välimaa t, törmälä p and tammela tlj (2002a), Interstitial laser coagulation and biodegradable self-expandable self-reinforced poly-L-lactic and poly-L-glycolic copolymer spiral stent in the treatment of benign prostatic enlargement, J Endourol, 16, 311–15. laaksovirta s, isotalo t, välimaa t, törmälä p, tammela tlj and talja m (2002b), Encrustation of biodegradable self-reinforced PLGA 80/20 copolymer, Memokath and Prostakath urethral stents, J Urol, 167 (suppl.), 363.

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laaksovirta s, laurila m, isotalo t, välimaa t, tammela tlj, törmälä p and talja m (2002c), Rabbit muscle and urethral in situ biocompatibility properties of the self-reinforced L-lactide-glycolic acid copolymer 80/20 spiral stent, J Urol, 167, 1527–31. madersbacher s (2005), Stents for prostatic diseases: Any progression after 25 years? Eur Urol, 49, 212–14. madersbacher s, alivizatos g, nordling j, sanz cr, emberton m and de la rosette jjmch (2004), EAU 2004 guidelines on assessment, therapy and follow-up of men with lower urinary tract symptoms suggestive of benign prostatic obstruction (BPH guidelines), Eur Urol, 46, 547–54. mcneill i and leiper h (1985), Degradation studies of some polyesters and polycarbonates – 2. Polylactide: degradation under isothermal conditions, thermal degradation mechanism and photolysis of polymer, Polym Degrad Stab, 11, 309–26. mehta ss and tophill pr (2006), Memokath® stents for the treatment of detrusor sphincter dyssynergia (DSD) in men with spinal cord injury: The Princess Royal Spinal Injuries Unit 10-year experience, Spinal Cord, 44, 1–6. multanen m, talja m, hallanvuo s, siitonen a, välimaa t, tammela tlj, seppälä and törmälä p (2000), Bacterial adherence to ofloxacin blended polylactonecoated self-reinforced L-lactic acid polymer urological stents, BJU Int, 86, 966–9. niesel t, moore rg, alfert hj and kavoussi lr (1995), Alternative endoscopic management in the treatment of urethral strictures, J Endourol, 154, 1117–18. nissenkorn i (1995), A simple nonmetal stent for treatment of urethral strictures: a preliminary report, J Urol, 154, 1117–18. nissenkorn i and shalev m (1997), Polyurethane stent for treatment of urethral strictures, J Endourol, 11, 481–3. oesterlink w and talja m (2000), Endoscopic urethroplasty with a free graft on a biodegradable polyglycolic acid spiral stent. A new technique, Eur Urol, 37, 112–15. perry mja, roodhouse ab, gidlow ab, spicer tg and ellis bw (2002), Thermoexpandable intraprostatic stent in bladder outlet obstruction: an 8-year stuty, BJU Int, 90, 216–21. petas a, talja m, tammela t, taari k, välimaa t and törmälä p (1997a), The biodegradable self-reinforced poly-L-lactic acid spiral stent compared with a suprapubic catheter in the treatment of post-operative urinary retention after visual laser ablation of the prostate, Br J Urol, 80, 439–43. petas a, kärkkäinen p, talja m, taari k, laato m, välimaa t and törmälä p (1997b), Effects of biodegradable self-reinforced polyglycolic acid, poly-DL-lactic acid and stainless steel spiral stents on uroepithelium after Nd:YAG laser irradiation of the canine prostate, Br J Urol, 80, 903–7. petas a, talja m, tammela t, taari k, lehtoranta k, välimaa t and törmälä p (1997c), A randomized study to compare biodegradable self-reinforced polyglycolic acid spiral stents to suprapubic and indwelling catheters after visual laser ablation of the prostate, J Urol, 157, 173–6. petas a, vuopio-varkila j, siitonen a, välimaa t, talja m and taari k (1998), Bacterial adherence to self-reinforced polylactic acid 96 spiral stents in vitro, Bioamaterials, 19, 677–81. petas a, isotalo t, talja m, tammela tl, välimaa t and törmälä p (2000), A randomised study to evaluate the efficacy of a biodegradable stent in the prevention

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of postoperative urinary retention after interstitial laser coagulation of the prostate, Scand J Urol Nephrol, 34, 262–6. shin jh, song hy, park h, kim jh, ko hk, kim yj, woo cw, kim th, ko gy, yoon hk and sung kb (2006), Removal of retrievable self-expandable urethral stents: experience in 58 stents, Eur Radiol, 16, 2037–43. song hy, park h, suh ts, gy ko, kim th, kim es and park t (2003), Recurrent traumatic urethral strictures near the external sphincter: treatment with a covered, retrievable, expandable nitinol stent – initial results, Radiology, 226, 433–40. stone ar, randall jr, shorrock k, peeling wb, rose mb and stephenson tp (1983), Optical urethrotomy in a 3 years experience, Br J Urol, 55, 701–4. steenkamp jw, heyns cf and dekock mls (1997), Internal urethrotomy versus dilatation as treatment for male urethral strictures: a prospective, randomized comparison, J Urol, 157, 98–101. terk md, stock rg and stone nn (1998), Identification of patients at increased risk for prolonged urinary retention following radioactive seed implantation of the prostate, J Urol, 160, 1379–82. törmälä p (1992), Biodegradable self-reinforced composite materials; manufacturing structure and mechanical properties, Clin Mater, 10, 29–34. törmälä p, pohjonen t and rikkanen p (1998), Bioabsorbable polymers: materials, technology and surgical applications, Proc Inst Mech Eng, 212, 101–7. vaajanen a, nuutinen j-p, isotalo t, törmälä p, tammela tlj and talja m (2003), Expansion and fixation properties of a new braided biodegradable uerthral stent: an experimental study in the rabbit, J Urol, 169, 1171–4. vainionpää s, rokkanen p and törmälä p (1989), Surgical applications of biodegradable polymers in human tissues, Prog Polym Sci, 14, 679–716. välimaa t and törmälä p (1996), Mechanical properties of bioabsorbable stents, Med Biol Eng Comput, 34 (suppl. I, part I), 20. välimaa t, talja m, tammela t, isotalo t, petas a, taari k and törmälä p (1998), Degradation properties of bioabsorbable self-reinforced self-expanding urological stents, J Endourol, 12, 6–19. välimaa t, laaksovirta s, tammela tlj, laippala p, talja m, isotalo t, petas a, taari k and törmälä p (2002), Viscoelastic memory and self-expansion of selfreinforced bioabsorbable stents, Biomaterials, 23, 3575–82. van dijk mm, mochtar ca, wijkstra h, laguna mp and de la rosette jjmch (2005), The bell-shaped nitinol prostatic stent in the treatement of lower urinary tract symptoms: experience in 108 patients, Eur Urol, 49, 353–9. vert m, li s and garreau h (1992), New insights on the degradation of bioabsorbable polymeric devices based on lactic and glycolic acids, Clin Mater, 10, 3–8. yachia d and beyar m (1991), Temporarily implanted coil stent for the treatment of recurrent urethral strictures: a preliminary report, J Urol, 146, 1001–4. yachia d (1997), Temporary metal stents in bladder outflow obstruction, J Endourol, 11, 459–65. yachia d and beyar m (1993), New self-expanding, self-retaining temporary coil stent for recurrent urethral stricture near the external sphincter, Br J Urol, 71, 317–21. yelderman jj and weaver rg (1967), The behavior and treatment of urethral strictures, J Urol, 97, 1040–4.

10 Penile implants G. B R O C K, University of Western Ontario, Canada

Abstract: Penile prosthetics have evolved both in materials and design since the initial devices were developed more than three decades ago. While oral medications for the management of erectile dysfunction (ED) are highly effective, roughly 30% of all men with ED fail to respond, have contra-indications to their use or have unacceptable side effects from these agents. The major challenges faced by clinicians implanting penile prosthetics are device malfunction, infection and poor functional results. Over the past decade, dramatic advances in materials used for penile prostheses have reduced infection rates by 50%, enhanced durability and provided improved cosmesis. In this chapter, these advances will be highlighted and critically reviewed. The potential for more physiologic devices through use of innovative biomaterials and device design are discussed; targeting novel cylinder, reservoir and valve pump designs. Key words: penile prosthesis, silicone, erectile dysfunction, inhibizone, bioflex.

10.1

Introduction

Great strides have been made in the field of penile prosthetics over the past decade.1–5 In spite of effective oral therapy for the management of erectile dysfunction (ED), penile implants remain a frequent therapeutic choice for tens of thousands of men annually refractory to pharmacotherapy.6,7 Looking down the research pipeline, innovative new molecules currently under development for ED include selective dopamine, glutamate, serotonin and melanocortin receptor agonists, rho-kinase inhibitors and guanylate cyclase activators; however, it seems unlikely these new agents will replace the need for surgical management of ED in all cases.8 The implant devices currently commercially available have a composite of enviable product characteristics, which allow the user to achieve near-normal physiologic erections while maintaining reasonable, concealability when not in use.9 While clearly inferior to a normal natural erection, contemporary penile prostheses have the ability to inflate and deflate within seconds, sustain enormous axial and radial compressive forces and remain pliable 226

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providing an acceptable cosmetic and functional result for an extended period of time over many years.10 Recent advances in the biomaterials and coatings used in penile prosthetics have provided additional durability and infection resistance, two aspects of these devices that have improved dramatically over the past decade, but which still remain a clinical area of concern.11–13 The contemporary design of an inflatable penile prosthesis is quite complicated. Its essential components consist of three elements: intra-penile paired cylinders, a valve pump that transfers fluid between components and a fluid reservoir. The cylinders, which are connected hydraulically, act as a single inflatable unit, able to provide complete rigidity and also deflate for total flaccidity. Cylinder design has undergone extensive trials and iterations over the past three decades. Currently, cylinders are available through a wide range of sizes, with normal girth and narrow girth and, in selected models, with the ability to expand in size with inflation.14,15 Transfer of the saline or contrast fluid is generally achieved through kink-resistant tubing of adequate luminal size to permit rapid inflation and deflation cycles. An early source of device failure was insecure connections between the tubing and reservoir and cylinders. This concern has been effectively overcome with snap connections and the elimination of all but a single connection in most devices (between the reservoir and the valve pump).16 The valve pump is the most complicated of all the components, yet needs to be sturdy enough to sustain high pressures when in use, while being unobtrusive and small in size. Finally the reservoir needs to have an adequate capacity and be able to transfer fluid to the device in a timely fashion, but not erode or lead to inflammation of adjacent organs. The basic design of the penile implant has not changed dramatically over its 35-year history. The biomaterials of which these devices are composed and the specialized coatings that are applied to implant surfaces, however, have revolutionized this industry and represent important clinical advances in the field.17 In this chapter, recent advances in biomaterials currently used for penile prosthetics will be highlighted with insight provided into what the next generation of materials and devices will hope to achieve.

10.2

Historical aspects of penile prosthesis development

In the earliest days, the idea of penile prosthetics was derived from preliminary designs for urinary incontinence devices developed at the University of Minnesota between 1968 and 1970. In late 1972, a newly formed company (American Medical Systems, AMS) retrofitted an artificial urinary cuff from

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1972

2007

10.1 The advances in design and materials used for inflatable penile implants over the past 35 years. The development of improved valve pumps, antibiotic coatings and improved cylinders are shown.

a sphincter device and replaced it with penile cylinders.18 These were silicone sheets reinforced with Dacron polyester fabric that produced non-expandable cylinders. In this earliest design, as shown in Fig. 10.1, two separate pumps were used for inflation and deflation. While this event marks a truly important time-point in penile prosthetics, the design requirements of a sphincter cuff and those of a penile implant cylinder are distinctly different. Penile cylinders, when in use, would need to sustain high axial loads, would be expected to be inflated to pressures exceeding three times systemic arterial pressure (300 mmHg) and be able to completely deflate. The sphincter cuff in contrast operates over a much lower pressure range, has a capacity of a much lower fluid volume and cycles from near capacity to near empty but is never truly exposed to either extreme.19,20 In this earliest device, a pressure-regulating valve controlled the fluid exchange from a pancake-type reservoir. The transfer of fluid was through medical grade tubing that was not kink resistant and manual pumping of the device was challenging.18 In total, just over a dozen devices of this type were ever implanted, but they represent a landmark in the evolution of the use of biomaterials and prosthetics in penile reconstructive surgery. Improvements in all device components have been achieved since that time, based on clinical experience and surgical outcomes.21–23 The basic concept of having an abdominally located reservoir, able to accommodate an ample amount of fluid connected via tubing (currently kink resistant and bacteria resistant) to two cylinders within the corpora cavernosa in which inflation

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and deflation cycles are controlled by a valve/pump mechanism located in the scrotum, remains intact today. Even the simplest component of these devices, the tubing, has specific design requirements: • • •

bend geometry of implant in vivo; comfort of patient – flex reduces pressure of tubing on tissue; cosmetics – flex allows for better conformance to tissue so does not protrude as noticeably.

The pump consists of two mechanisms: a bulb that is squeezed to push fluid from the reservoir into the cylinders, and a release mechanism that is squeezed to allow the fluid to flow back into the reservoir. Both of these mechanisms need to be activated with a reasonable amount of force. Based on studies of hand strength, which varies with advancing age and with squeeze method from 13 to 27 lbs (6–12 kg), a specification of pressure to induce pump activation at an appropriate level has been defined.24 Pumps fabricated out of more durable materials than silicone, required excessive force to activate the bulb and deflate. A range of cylinder lengths and widths is available as the cylinders in common use range from 12 to 25 cm in length with the option of adding rear tip extenders to further match measured intra-corporal length. Variation in fluid volumes transferred in each cycle of the device range from a few cc’s to nearly 100 cc for the largest devices, in which girth and length expansion are possible. The advances in materials, manufacturing and component design have allowed for enhanced reliability, durability and physiologic functioning. Ultimately, however, the degree of patient and partner satisfaction determine whether these devices can be considered as successful.25–29 In sharp contrast to these devices currently approved and marketed, the ideal penile prosthesis would exactly replicate the normal erection, characterized by a rapid erectile response within seconds, have the ability to retain the intra-cylinder pressures of 100–300 mmHg for minutes–hours, achieve complete detumescence with total flaccidity and provide exquisite sensitivity during all phases.30,31 While significant improvements have occurred in these devices over the past four decades, current devices still lack total concealability, remain palpable in most cases at the distal end beneath the glans and often have tubing and a valve pump that may be apparent. In addition, device malfunction, erosion, infection and discomfort associated with the materials used still remain an obstacle to total device satisfaction in a portion of those implanted.32,33 While the mechanical design of the device is important, the actual biomaterials chosen and their ability to meet a number of key material design criteria often define success. These characteristics include: fatigue and abrasion resistance, ease of processing, providing consistent material characteristics, reliable stability in vivo and the material must be fully biocompatible.

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One of the features that is most important to clinicians and users alike is the ability to resist infection.

10.3

Biomaterials in current use

Penile prostheses exist in a variety of forms and materials. The multicomponent inflatable devices are the most complex, with hydraulic tubing, cylinders, rear tip extenders, connection mechanisms and reservoirs being the key elements. The various components of the inflatable penile prosthesis have different design requirements and as a consequence are typically composed of different materials. In contrast, malleable prostheses are composed of durable inner core elements, providing rigidity with a modest amount of flexibility, surrounded by a biocompatible outer coating that is designed to resist infection or inducing host inflammation. The focus of this chapter is inflatable devices, since they represent the state of the art, provide the most physiologic result and are the most commonly implanted devices.

10.3.1 Design challenges of the inflatable penile prosthesis Several highly significant challenges exist in the development of an ideal penile prosthesis. In contrast to many prosthetic devices, the inflatable implant must be cycled through inflation and deflation cycles, and remain malleable and supple yet durable throughout a product life of decades. The list below highlights the specific design objectives: • • • • • • • • • •

to be biocompatible; to sustain multiple inflation/deflation cycles; to resist infection and biofilm formation; to resist device rupture, erosion, deformation and malfunction; to have easy and reliable tubing connections; to have a reservoir that can accommodate ample amounts of fluid; to have feel and function in that is natural to the host and partner; to be ergonomically designed to allow for normal activities of daily living; to be cost efficient and be manufactured on a large industrial scale; to inflate/deflate in a physiologically reasonable time frame and to a physiologically reasonable extent.

10.3.2 Cylinder materials Inflatable penile prosthesis cylinders are the core element of the device. They provide the functionality to the implant and define the extent to which

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the patient will be satisfied with their surgical result. The cylinders undergo significant stress during inflation and use (they will be subject to a combination of flex, tension, compressive buckling and torsional stresses). Abrasion can occur between the cylinder and device tubing, as well as between the device surface and the inner layer of the tunica albuginea where a capsule generally forms within weeks.34–37 Body fluids and immune systems will expose the device to chemical degradation. The components of the device will be stressed during sexual activity but will also sustain ongoing frequent strain during normal activities of daily living – such as walking and running – that involve flexion of the pelvis. Defined sites of mechanical failure have been identified at the junction of the tubing insertion into the cylinders, and at sites of tubing connections. The numbers of connections have been reduced from three to a single connection in an attempt to reduce intraoperative time and limit failure at these sites. Earlier versions of the cylinders sustained an unacceptably high rate of aneurismal formation caused by silicone failure. This problem is no longer commonly seen because of the use of reinforced materials and the use of a sandwich configuration, which allows a certain degree of material expansion and affords great axial rigidity.33,38

10.3.3 Reservoir materials The reservoir is an essential component of the inflatable penile implant, allowing complete detumescence of the cylinders when not in use and rapid transfer of fluid to the penis when required. The early approaches to this element of the three-piece penile implant involved the use of a ‘pancake’ reservoir, which was flat and Dacron reinforced.39 It rapidly became apparent that the ideal design characteristics for a reservoir would be: (a) zero pressure when filled; (b) able to sustain multiple cycles of filling and emptying; and (c) adequate capacity to completely empty and fill cylinders ranging in size from 12 to 25 cm, with a functional capacity of 60–100 cc. Subsequent modifications of the reservoir eliminated the reinforced Dacron layer with the addition of a dip-coated silicone and reinforced connecting anti-kink tubing. In the current versions of the reservoirs for both Coloplast and AMS devices, antibacterial surface coatings and enhanced resistance to erosion have been applied. Strategies to reduce the major complications associated with penile prostheses have focused on new materials and device coatings over the past decade; these have included surface coatings to enhance durability and reduce infections. Infections of the device represent the most dreaded frequent complication of the procedure and are a constant source of concern to clinicians.

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10.4

Device infection

Implanted prosthetic devices of all types are at risk of infection at the time of implantation because they provide a platform for the development of a bacterial biofilm, an organized bacterial colony that grows on the surface of the implanted material.40–42 Biofilm is a microbial-derived sessile community of cells that permanently attach to a material or to each other and become embedded in a matrix of extracellular polymeric substances that they have produced.43 The biofilm is typically resistant to all antibiotics and mechanical means of eradication, short of explantation of the prosthesis. Bacterial adherence and biofilm production proceed in two steps: first, an attachment to a surface and, second, a cell-to-cell adhesion. Recent reports have described a variety of strategies for bacterial attachment to the surface of the foreign material: (a) surface charge attraction; (b) hydrophilic/hydrophobic interactions; and (c) specific attachment by fimbrae.43–45 As with all bacterial infections, growth follows bacterial adhesion to the biofilm and is usually thought of as being composed of three distinct layers: (a) a binding layer of bacteria attached directly to the biofilm; (b) a compact layer of bacteria making up the basal layer; and (c) a superficial non-attached layer which can spread distantly.43,46 The clinical dilemma is that the biofilm provides an ideal environment for bacterial growth, in a sanctuary free from antibacterial agents, typically with a low perfusion environment and low oxygen concentrations further hindering antibiotics focusing on metabolic rates. Predominant among the various strains of bacteria causing penile implant infections are nosocomial staphylococcal organisms, which express a gene producing a biofilm phenotype characterized by a polysaccharide intercellular adhesin molecule.43 The pathogenicity of Staphylococcus epidermidis is mostly due to its ability to colonize indwelling polymeric devices and form a thick, multi-layered biofilm. Biofilms play an important role in the spread of antibiotic resistance. Within the dense bacterial population, efficient horizontal transfer of resistance and virulence genes takes place. AMS has chosen to apply an antibiotic coating of minocycline and rifampin, which they call inhibizone (Fig. 10.2). It has been shown to persist on the surface of the device and provide an area of bacterial inhibition for 7 days. In a recent report from Carson,11 4205 original implants were followed for 1–2 years. A total of 2261 devices were impregnated with inhibizone and these were compared with a contemporary series of 1944 non-inhibizone devices. Infection rates among the non-treated devices were significantly higher than the inhibizone cohort, with infection rates of 1.68% compared with 0.68% at 180 days, clearly demonstrating the clinical efficacy of this approach.11 These findings support earlier in vitro studies showing significant zones of inhibition surrounding inhibizone-coated tubing samples

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10.2 Inhibizone coating.

10.3 Hydrophilic coating.

of 0.6–2.2 cm for the strains of bacteria that most frequently cause device infections.47–49 Coloplast, the manufacturers of the Titan implants, have employed an alternative approach whereby a hydrophilic surface, called ‘resist’ (Fig. 10.3) has been applied to the devices. This surface reportedly both absorbs

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antibiotics when immersed prior to implantation and provides the device with a surface to which bacteria are less likely to adhere. The advantage of this approach is that the implanting surgeon has the ability to select the specific antibiotics in each case based on the bacterial flora in that geographical region and based on patient sensitivity. The hydrophilic surface coating also allows ease of placement of the cylinders, which slide into the corpora and compress to a greater degree than the silicone sandwich materials of the AMS devices; the hydrophilic coating therefore permits implantation through a slightly smaller corporotomy.50,51 The biofilm is generally thought to be composed of Staphylococcus aureus, Pseudomonas aeruginosa and S. epidermidis.

10.5

Erosion resistance

Parylene coating was introduced into the design of AMS cylinders in 2000 in an attempt to enhance device and tubing durability (Fig. 10.4). It has been proven to increase material fatigue resistance, limit the abrasive effects of tubing on the cylinders and is therefore a device enhancement.52 The basic material used for inflatable prostheses is silicone. It has excellent biocompatibility properties, but has demonstrated known problems with fatigue and limited abrasion resistance. Silicone can also be difficult to process and there can be high variability in material properties from batch to batch. A

10.4 Silicone sandwich with fabric and Parylene coating.

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Table 10.1 The relative strength and material characteristics of silicone and bioflex used in contemporary penile implants. While bioflex has enviable strength properties it is not suitable for all components

Property

Bioflex

Typical platinum cure medical grade silicone

Tensile strength Elongation Tear strength (die B) Tensile modulus at 200% elongation

7500 psi 600% 400+ ppi 500 psi

1450 psi 1000% 250 ppi 300–350 psi

ppi, pounds per inch; psi, pounds per square inch.

variety of approaches have been developed to overcome these issues and reduce device failure rates – including sandwich designs with layers of silicone covering a polyester fabric inner sheath, Parylene coating and reinforcing material cuffs. An alternative design material is bioflex.53,54 This is a urethane elastomer used by Coloplast for cylinders and reservoir bladders. Urethane elastomer has the potential to fulfill many of the material design criteria. Comparative testing between the typical platinum cure medical grade silicone and Bioflex is shown in Table 10.1. While the ability to elongate is greater for silicone, most contemporary devices are not designed to expand in girth or length. Among the key material characteristics shown in Table 10.1 are tensile strength and tear strength. These properties depict the likelihood of material failure. The strength and durability of bioflex allows its use in monolayer designs, which eliminates the abrasion inherent in multi-layer systems. The monolayer design also allows a simple molding process to be used for fabrication of the bladders.

10.6

Summary

Advances in materials and design have contributed to improved device satisfaction rates, enhanced performance characteristics of biocompatibility, reliability and comfort, and reduced infection rates. The field of penile prosthetics has evolved rapidly over the past decade, in part as a result of improved oral treatment alternatives for ED, patient expectations and significant dedicated biomaterials research. While the perfect device material is yet to be discovered, contemporary penile implants have demonstrated low device failure rates with high patient and partner satisfaction.

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10.7

Future trends

The past few decades have been witness to the rapid development of commercially viable penile prosthetic devices, which attempt to simulate a normal natural erection. Early devices were fraught with mechanical malfunctions, erosions and infections, and had relatively poor tumescence and detumescence properties. The use of a silicone-based device was in retrospect the first large leap forward towards a physiologically acceptable implant and over the past decade the use of new materials, as outlined in this chapter (such as the addition of surface coatings able to reduce infection rates by 50%, the enhanced durability of the cylinders and improved tubing connections) has led to an enviable reliability rate. In spite of these many improvements over the years, currently available penile implants still do not function as well as a normal natural erection. Patients frequently encounter difficulty in the inflation and detumescence of the implant and instructing patients on the use of their devices is frequently a challenging and time-consuming exercise for clinicians. The relatively large size of the scrotal valve-pump can be a source of patient discomfort and dissatisfaction. Redesigning of the pump to produce an electronically controlled miniature device would seem like a logical next step, in concert with cylinders that are bacteriostatic, resist erosion of the host tissues, and remain soft, pliable and durable over extended periods of time. Currently, implantation of inflatable devices requires placement of a retropubically located reservoir. In cases where radical pelvic surgery has been performed (such as radical cystectomy) this space may be obliterated and it can become a surgical challenge to develop adequate space for the 60–100 cc reservoirs. The ability to place a pancake or flat reservoir under the abdominal musculature would avoid the need for a pelvic dissection and may represent an advance in therapy in the near future. Lock-out valves preventing auto-inflation of the cylinders has been introduced over the past 5 years and have generally been perceived as an advance by most clinicians and patients. Finally, one area of research that has been explored over the past decade but has yet to be successfully resolved is that of girth and length enhancement; this may prove to be the ultimate advance in penile prosthetics. Among patients with Peyronie’s disease or post-radical pelvic surgery, significant length loss results from their disease process or surgical intervention. The majority of men undergoing surgical penile prosthesis placement will experience a significant amount of penile length and girth loss from the procedure itself. The ability of future devices to compensate and correct for this, with the potential for actual augmentation, would be a truly significant advance in prosthetics.

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The ideal device of the future would be composed of a biomaterial able to resist infection, durable enough to sustain hundreds of thousands of cycles, rigid yet expandable and able to train the tunica albuginea and corporal tissues into greater length and widths over time. The mechanical components would be made of more pliable materials, less obtrusive to the user and partner, yet retain the robust design criteria allowing many years of use with electronics to allow for size reduction.

10.8

References

1 henry gd, wilson sk. Updates in inflatable penile prostheses. Urol Clin North Am 2007;34(4):535–47. 2 mulcahy jj, wilson sk. Current use of penile implants in erectile dysfunction. Curr Urol Rep 2006;7:485–9. 3 wilson sk, delk jr 2nd. Historical advances in penile prostheses. Int J Impot Res 2000;12(suppl. 4):S101–7. 4 simmons m, montague dk. Penile prosthesis implantation: past, present and future. Int J Impot Res 2008;20(5):437–44. 5 natali a, olianas r, fisch m. Penile implantation in Europe: successes and complications with 253 implants in Italy and Germany. J Sex Med 2008;5(6):1503–12. 6 shabsigh r, duval s, shah m, regan ts, juhasz m, veltry lg. Efficacy of vardenafil for the treatment of erectile dysfunction in men with hypertension: a meta-analysis of clinical trial data. Curr Med Res Opin 2007;23(10): 2453–60. 7 hatzimouratidis k, hatzichristou dg. A comparative review of the options for treatment of erectile dysfunction: which treatment for which patient? Drugs 2005;65(12):1621–50. 8 hatzimouratidis k, hatzichristou dg. Looking to the future for erectile dysfunction therapies. Drugs 2008;68(2):231–50. 9 khoudary kp, morgentaler a. Design considerations in penile prostheses: the American Medical Systems product line. J Long Term Eff Med Implants 1997;7(1):55–64. 10 levine la, estrada cr, morgentaler a. Mechanical reliability and safety of, and patient satisfaction with the Ambicor inflatable penile prosthesis: Results of a 2 center study. J Urol 2001;166:932–7. 11 carson cc 3rd. Efficacy of antibiotic impregnation of inflatable penile prostheses in decreasing infection in original implants. J Urol 2004;171:1611–14. 12 wolter ce, hellstrom wj. The hydrophilic-coated inflatable penile prosthesis: 1-year experience. J Sex Med 2004;1:221–4. 13 delk j, knoll ld, mcmurray j, shore n, wilson s. Early experience with the American Medical Systems new tactile pump: Results of a multicenter study. J Sex Med 2005;2:266–71. 14 milbank aj, montague dk, angermeier kw, lakin mm, worley se. Mechanical failure of the American Medical Systems Ultrex inflatable penile prosthesis: before and after 1993 structural modification. J Urol 2002;167(6):2502–6.

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Biomaterials and tissue engineering in urology deuk choi y, jin choi y, hwan kim j, ki choi h. Mechanical reliability of the AMS 700CXM inflatable penile prosthesis for the treatment of male erectile dysfunction. J Urol 2001;165(3):822–4. montague dk, lakin mm. Early experience with the controlled girth and length expanding cylinder of the American Medical Systems Ultrex penile prosthesis. J Urol 1992;148:1444–6. dhar nb, angermeier kw, montague dk. Long-term mechanical reliability of AMS 700CX/CXM inflatable penile prosthesis. J Urol 2006;176:2599–601. furlow wl. Surgical management of impotence using the inflatable penile prosthesis. Experience with 36 patients. Mayo Clin Proc 1976;51(6):325–8. scott fb, bradley we, timm gw. Management of erectile impotence: use of implantable inflatable prosthesis. Urology 1973;2:80–2. mohammed a, khan a, shaikh t, shergill is, junaid i. The artificial urinary sphincter. Expert Rev Med Devices 2007;4(4):567–75. sadeghi-nejad h, sharma a, irwin rj, wilson sk, delk jr. Reservoir herniation as a complication of three-piece penile prosthesis insertion. Urology 2001;57:142–5. cuellar dc, sklar gn. Penile prosthesis in the organ transplant recipient. Urology 2001;57:138–41. lockyer r, gingell c. Spontaneous breakage of malleable prosthesis. Int J Impot Res 1999;11:237. wilson sk, henry gd, delk jr jr, cleves ma. The mentor Alpha 1 penile prosthesis with reservoir lock-out valve: Effective prevention of auto-inflation with improved capability for ectopic reservoir placement. J Urol 2002; 168(4 Pt 1):1475–8. mathiowetz v, kashman n, volland g, weber k, dowe m, rogers s. Grip and pinch strength: normative data for adults. Arch Phys Med Rehabil 1985;66:69–72. daitch ja, angermeier kw, lakin mm, ingleright bj, montague dk. Long-term mechanical reliability of AMS 700 series inflatable penile prostheses: Comparison of CX/CXM and Ultrex cylinders. J Urol 1997;158:1400–2. dubocq f, tefilli mv, gheiler el, li h, dhabuwala cb. Long-term mechanical reliability of multicomponent inflatable penile prosthesis: comparison of device survival. Urology 1998;52:277–81. akin-olugbade o, parker m, guhring p, mulhall j. Determinants of patient satisfaction following penile prosthesis surgery. J Sex Med 2006;3: 743–8. levine la, estrada cr, morgentaler a. Mechanical reliability and safety of, and patient satisfaction with the Ambicor inflatable penile prosthesis: Results of a 2 center study. J Urol 2001;166:932–7. wilson sk, cleves ma, delk jr 2nd. Ultrex cylinders: Problems with uncontrolled lengthening (the S-shaped deformity). J Urol 1996;155:135–7. lue tf, takamura t, schmidt ra, palubinskas aj, tanagho ea Hemodynamics of erection in the monkey. J Urol 1983;130(6):1237–41. lue tf, takamura t, umraiya m, schmidt ra, tanagho ea Hemodynamics of canine corpora cavernosa during erection. Urology 1984;24(4):347–52. small mp, carrion hm, gordon ja. Small-Carrion penile prosthesis. New implant for management of impotence. Urology 1975;5:479–86.

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carson cc, mulcahy jj, govier fe. Efficacy, safety and patient satisfaction outcomes of the AMS 700CX inflatable penile prosthesis: Results of a long-term multicenter study. AMS 700CX Study Group. J Urol 2000;164:376–80. Ferguson KH, Cespedes RD. Prospective long-term results and quality-of-life assessment after Dura-II penile prosthesis placement. Urology 2003;61:437– 441. goldstein i, bertero eb, kaufman jm, witten fr, hubbard jg, fitch wp et al. Early experience with the first pre-connected 3-piece inflatable penile prosthesis: the Mentor Alpha-1. [see comments]. J Urol 1993;150: 1814–8. small mp. Small–Carrion penile prosthesis: a report on 160 cases and review of the literature. J Urol 1978;119:365–8. carson c. Complications of penile prostheses and complex implantations. In: Carson C, Kirby R, Goldstein I, eds. Textbook of Male Erectile Dysfunction. Oxford, UK: Isis Medical Media; 1999; pp. 435–50. goldstein i, newman l, baum n, brooks m, chaikin l, goldberg k, mcbride a, krane rj. Safety and efficacy outcome of mentor alpha-1 inflatable penile prosthesis implantation for impotence treatment. J Urol 1997;157: 833–9. furlow wl. Inflatable penile prosthesis: Mayo clinic experience with 175 patients. Urology 1979;13:166–71. carson cc. Infections in genitourinary prostheses. Urol Clin North Am 1989;16:139–47. darouiche ro. Treatment of infections associated with surgical implants. N Engl J Med 2004;350:1422–9. sadeghi-nejad h. Penile prosthesis surgery: A review of prosthetic devices and associated complications. J Sex Med 2007;4:296–309. silverstein a, donatucci cf. Bacterial biofilms and implantable prosthetic devices. Int J Impot Res 2003;15(suppl. 5):S150–4. jarow jp. Risk factors for penile prosthetic infection. J Urol 1996;156(2 Pt 1):402–4. wilson sk, delk jr 2nd. Inflatable penile implant infection: Predisposing factors and treatment suggestions. J Urol 1995;153(3 Pt 1):659–61. montague dk, angermeier kw, lakin mm. Penile prosthesis infections. Int J Impot Res 2001;13:326–8. wilson sk, carson cc, cleves ma, delk jr. Quantifying risk of penile prostheses infection with elevated glycosylated hemoglobin. J Urol 1998;159: 1537–40. wilson sk, zumbe j, henry gd, salem ea, delk jr, cleves ma. Infection reduction using antibiotic-coated inflatable penile prosthesis. Urology 2007;70(2): 337–40. abouassaly r, angermeier kw, montague dk. Risk of infection with an antibiotic coated penile prosthesis at device replacement for mechanical failure. J Urol 2006;176(6 Pt 1):2471–3. droggin d, shabsigh r, anastasiadis ag. Antibiotic coating reduces penile prosthesis infection. J Sex Med 2005;2:565–8. wolter ce, hellstrom wj. The hydrophilic-coated inflatable penile prosthesis: 1-year experience. J Sex Med 2004;1:221–4.

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wilson sk, delk jr, salem ea, cleves ma. Long-term survival of inflatable penile prostheses: single surgical group experience with 2,384 first-time implants spanning two decades. J Sex Med 2007;4(4 Pt 1):1074–9. 53 rajpurkar a, fairfax m, li h, dhabuwala cb. Antibiotic soaked hydrophilic coated bioflex: a new strategy in the prevention of penile prosthesis infection. J Sex Med 2004;1(2):215–20. 54 hellstrom wj, hyun js, human l, sanabria ja, bivalacqua tj, leungwattanakij s. Antimicrobial activity of antibiotic-soaked, Resist-coated Bioflex. Int J Impot Res 2003;15(1):18–21.

11 Artificial biomaterials for urological tissue engineering W. A. FA R H AT, The Hospital for Sick Children, Canada; and P. J. G E U T J E S, Radboud University Nijmegen Medical Centre, The Netherlands

Abstract: Formation of new biological tissues by tissue engineering is achieved in vitro and in vivo by cells that are regulated by certain growth factors/cytokines and by matrices with specific structures. The extracellular matrix (ECM) is a complex ordered structure that contains multiple components (collagens, proteoglycans, structural glycoproteins and elastin) of macromolecules. The composition of the ECM is not static; rather it changes during normal development and tissue repair and regeneration. The ECM is the optimized structural milieu that maintains cell homeostasis and directs tissue development. Tissue engineering generally requires an artificial ECM for tissue regeneration. Otherwise, cell proliferation and differentiation, resulting in tissue regeneration, would be difficult unless such a matrix is provided that functions as a natural ECM. Placed at the site of a defect, such materials should actively and temporarily participate in the regeneration process by providing a platform on which cell-triggered remodeling could occur. Furthermore, since this artificial ECM should disappear through absorption into the body after implantation and when the new tissue is regenerated, it is preferable that materials for the matrix are prepared from biodegradable polymers to assist tissue regeneration. Key words: tissue engineering, biomaterials, urology.

11.1

Introduction

The ECM is a complex of macromolecules capable of self-assembly, composed predominantly of collagens, glycoproteins and proteoglycans. The native ECM in the body is a complex and dynamic environment filled with nanofeatures of pores and fibers that exhibit tissue-specific structure and properties. Therefore, the characteristics of a scaffold vary according to the tissue types where the scaffold is to be used. This matrix acts as a scaffold for cells, serves as a reservoir for growth factors and modulates their activation status, and therefore profoundly influences the biological behavior of cells in terms of cell growth, differentiation, 243

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development and metabolic responses.1–3 Therefore, a great effort has been made to mimic the ECM to guide morphogenesis in tissue repair and tissue engineering.4–6 The artificial scaffold, by definition, is a temporary supporting structure for growing cells and tissues. It is also called synthetic ECM and plays a critical role in supporting the cells to accommodate the new environment. These cells then undergo proliferation, migration and differentiation in three dimensions, which eventually leads to the formation of a specific tissue with appropriate functions as would be found in the human body. In order to facilitate these measures, the scaffold should possess a few basic characteristics. An ideal scaffold for tissue engineering should possess all the qualities of a native ECM and should function in the same way as the ECM under physiological conditions. Furthermore, these matrices must display some key characteristics of the ECM of the organ interest. The synthetic ECM should provide temporary mechanical support sufficient to withstand in vivo forces and maintain a potential space for tissue development. This mechanical support by the synthetic ECM should be maintained until the engineered tissue has sufficient mechanical integrity to support itself. Although mimicking the complexity of the native ECM is not so easy, recent investigations suggest that the scaffold for tissue engineering should also have some of the basic characteristics essential for tissue development,7 such as induction and preservation of the appropriate cell phenotype and gene expression.

11.2

History of synthetic biomaterials used in urology

Synthetic biomaterials exhibit several attractive features for applications in tissue engineering. Two benefit shared by all synthetic biomaterials is that they minimize the risk of carrying biological pathogens or contaminants and that they may be designed with controlled drug release that results in tissue repair and tissue engineering.8 Recently developed synthetic biomaterials show promising improvements in in vivo biocompatibility. The major benefit of synthetic biomaterials is that they can be designed to meet specific needs. In some cases, synthetic biomaterials are composed of polymers of naturally occurring small biological molecules such as amino acids. An example of such design may be achieved by the incorporation of biologically active motifs that promote cell attachment (e.g. the cell adhesion motif arginine, glycine, aspartic acid (RGD), a ligand for integrin cell adhesion receptors). The basic units of synthetic biomaterials show excellent physiological compatibility and minimal cytotoxicity, and the breakdown products of biomaterials that are derived from biological molecules can be incorporated into newly synthesized biomolecules or metabolized in the host organism.

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Generally, the advantages of US Food and Drug Administration (FDA)approved synthetic polymers such as polyesters of naturally occurring α-hydroxy acids – including polyglycolic acid (PGA), polylactic acid (PLA) and poly(lactic-co-glycolic acid) (PLGA) – make their use in the engineering of genitourinary tissues common. The synthetic polymers can be manufactured with consistent reproducibility on a large scale with well-controlled strength, degradation rate and microstructure properties. The ester bonds in these polymers are hydrolytically labile, and the polymer degradation products are non-toxic, natural metabolites that are eventually eliminated from the body in the form of carbon dioxide and water.9 The degradation rate of these polymers can be tailored from several weeks to several years by alteration of their crystallinity, initial molecular weight and the copolymer ratio of lactic to glycolic acid. Furthermore, using various techniques, these polymers can easily be formed into three-dimensional scaffolds with the desired microstructure, gross shape and dimensions. Furthermore, the mechanical properties of the scaffold can be controlled by the fabrication process. Many applications in genitourinary-tissue engineering often require a scaffold with a high degree of porosity and a high surface-areato-volume ratio. This has been addressed by the processing of biomaterials into configurations of fiber meshes and porous sponges using the techniques described above.

11.3

Synthetic scaffolds

11.3.1 General considerations For tissue engineering purposes, unique scaffold characteristics are mandatory. Scaffolds are implanted as intact grafts following processing, into which host cells grow, proliferate and differentiate. Thus, this approach places a high priority on the native structure and composition of the synthetic scaffold and its inherent ability to interact with the host. Scaffolds or biomaterials are expected to function as an artificial ECM to facilitate the formation of new tissues with appropriate structure and with the appropriate biological and mechanical functions of native ECM. The selected scaffold should be biodegradable and bioresorbable to support the reconstruction of a completely normal tissue without inflammation. Such behavior of the scaffolds would avoid the risk of inflammatory or foreign-body responses that may be associated with the permanent presence of a foreign material in vivo. The degradation products should not provoke inflammation or toxicity and must be removed from the body via metabolic pathways. The degradation rate and the concentration of degradation products in the tissues surrounding the implant must be at tolerable levels. Hence, implantable scaffolds designed to replace damaged or diseased organs must act as

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supports into which cells can migrate, must be strong enough to withstand the physiological and mechanical properties demands placed upon them and must establish the required blood supply. In order to achieve an uncomplicated outcome after biomedical application of scaffolds, successful wound healing is a must.

11.3.2 Specific considerations Scaffolds need to be processed into specific configurations according to the organ of interest. For instance, a large surface-area-to-volume ratio is often desirable in order to allow the delivery of a high density of cells. Furthermore, the configuration of the biomaterials can guide the formation and structure of an engineered tissue; highly porous scaffolds are desirable in order to allow cell migration which in turn affects tissue in-growth and cellular repopulation. In addition, the scaffolds provide mechanical support against in vivo forces, thus maintaining a predefined structure during the process of tissue development. The scaffolds provide temporary mechanical support sufficient to withstand in vivo forces exerted by the surrounding tissue and maintain a potential space for tissue development. The mechanical support of the scaffolds should be maintained until the engineered tissue has sufficient mechanical integrity to support itself. This can potentially be achieved by an appropriate choice of mechanical and degradative properties of the biomaterials. Hence, the scaffold must be capable of controlling the structure and function of the engineered tissue in a pre-designed manner by interacting with transplanted cells and/or the host cells. Furthermore, the synthetic scaffolds can be further modified with bioactive signals, and can also serve as a depot for the local release of growth factors and other bioactive agents that induce tissue-specific gene expression of the cells which can regulate cellular function. The architecture of scaffolds is believed to contribute significantly to the development of engineered tissues and is thought to provide appropriate nutritional conditions and spatial organization for cell growth, with porosity being an important physical property that ultimately determines the architecture of the scaffolds. For instance, a highly porous scaffold is required to accommodate mammalian cells and guides their growth and tissue regeneration in three dimensions. Furthermore, highly porous biomaterials are desirable for the easy diffusion of nutrients to, and waste products from, the implant; most importantly, the pores allow vascularization to occur more quickly. On the other hand, each tissue or organ has its own characteristic architectural organization, which is closely related to its function. Normally, leakage of urine from the bladder via the wall does not occur; therefore regenerated bladder substitutes should not allow leakage even after

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longstanding urine retention. Hence, in a urinary bladder replacement setting, porous scaffolds may have an impact on urine leakage, resulting in adverse outcome after bladder substitution. The permeability of tissue could be affected by the processing techniques used, and thus impact nutrient flow and leakage, particularly in the absence of urinary diversion. Hence it is essential that urinary bladder cells are seeded on to the synthetic scaffolds before implantation in order to enhance their regenerative capacity.

11.4

Smart biomaterials

The greatest disadvantage with traditional synthetic materials, however, is the lack of cell-recognition signals that regulate cell adhesion functions. Rather, the synthetic scaffolds promote cell adhesion via indirect recognition, i.e. by proteins (e.g. fibronectin and vitronectin) from the body fluids adsorbing non-specifically to the material surface. Since the design and selection of the scaffold is critical in the development of engineered tissues, appropriate regulation of cell behavior – such as adhesion, proliferation, migration and differentiation – are mandatory requirements in the design of the artificial or synthetic scaffolds. More importantly, the scaffold should have a suitable architecture to promote cellular interaction and tissue development, and possess proper mechanical and physical properties, while having reduced resistance to urine. Failure of cellular repopulation of the synthetic scaffolds makes them susceptible to loss of structural integrity secondary to inflammation and fibrosis caused by urine. Furthermore, evidence in the field now supports the idea that attaching a vascular supply to the scaffold may be necessary for reconstruction of bladder defects. Permeability to urine and lack of prompt blood supply to the scaffolds leads to the dysregulation of these processes. Combining the advantages of both synthetic materials and naturally derived materials is an emerging area and is an attractive option in tissue engineering. A new innovative direct approach to promote the bioavailability of scaffolds would be the incorporation of cell adhesion peptides, growth factors and DNA found in natural ECM into synthetic materials. The hybrid materials would possess the specific biological activities of naturally derived materials as well as the favorable properties of synthetic materials, including widely controllable mechanical properties and good processability. On the other hand, in order to design a scaffold capable of inducing tissue-specific gene expression of cells, it would be necessary to identify the key molecular components involved in the cell behaviors required for regulation of tissue function. For instance, the physicochemical principles that influence receptor-mediated cell regulation, such as receptor–ligand interactions, have to be addressed. In addition, since organs are usually present in a dynamic state, the expression of genes by cells in engineered tissues may also be regulated by

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interactions with other cells, growth factors and mechanical stimuli imposed on the cells. Another well-described approach within tissue engineering is to construct molecularly defined bioscaffolds to form the desired tissue.10–12 These so-called ‘smart scaffolds’ should not only resemble the ECM in vivo but also create the molecular microenvironment to induce and sustain growth, migration and differentiation of organ-specific types of cells. When cells are seeded on non-specific scaffold, they will dedifferentiate and thus will not be able to form organ-specific tissues. Thus, to design and construct smart and tailor-made scaffolds, a number of aspects have to be taken into account such as biochemical composition, growth factor or cytokine gradient, microstructure and mechanical strength. In addition, the construct should provide temporary mechanical support sufficient to withstand in vivo forces and maintain structure for tissue cell in-growth. Several basic tailor-made scaffolds have been prepared mimicking the desired microenvironment and architecture with an appropriate porous morphology, using highly purified biomaterials including specific ECM molecules (collagen, fibronectin, elastin and/or laminin), modifying molecules (glycosaminoglycans and/or integrins) and effector molecules (growth factors and/or cytokines).13–15 Such scaffolds have predefined physicochemical, biomechanical and morphological characteristics, and – especially when combined with growth factors – show great potential for tissue engineering of, for example, bladder, urethra and ureter.16,17

11.4.1 Components of smart biomaterials In order to act as a template for tissue regeneration and be integrated in normal biological function and healing, synthetic scaffolds should contain the major constituents of the dynamic ECM. Collagen Collagen is the most widely used ECM protein in urological tissue engineering.18–20 Type I collagen is the major collagen in all organs, especially those requiring strength and stability like tendon, skin and bone. The basic structural collagen, tropocollagen, has a characteristic helical structure (1.5 × 300 nm) consisting of three polypeptide α-chains which are coiled to form a right-handed triple helix. Collagen fibers are formed by the assembly of lateral and end-to-end aggregation of microfibrils. Collagen has a variety of functional properties favorable for cellular growth. It is a biocompatible, biodegradable and non-toxic polymer. Collagen mediates biological functions such as the binding, migration, growth and chemotaxis of cells, and interactions with matrix proteins and

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proteoglycans. Insoluble type I collagen has found ample usage in biomedical fields.21–26 However, collagen extracted from natural tissues is capable of eliciting certain immunogenic responses upon implantation; thereby, direct use of this type of collagen is limited. Nowadays, a purified form of collagen, known as ‘reconstituted collagen’, is produced by biochemical processing; this elicits relatively less immunogenic response and can therefore be used for tissue engineering applications. It provides strength and structural integrity to tissues like skin, cartilage, bone, tendon and dental elements. Hence, collagen-based constructs need stable chemical cross-linking to control the mechanical properties and the degradation time, and to some extent the immunogenicity of the device. Elastin Elastin and elastin-derived peptides that confer flexibility and distensibility to all tissues have been combined with various biological matrices to modulate their morphological, physical and biological characteristics.10 Elastin contains hydrophobic amino acids (proline, glycine, desmosine and isodesmosine) and a high degree of intermolecular cross-links which makes elastin fibers highly resistant to proteolytic degradation. Elastin fibers are able to recoil after stretching; furthermore, the long-term stability of the elastin fibers makes it a desirable protein for dynamic organ tissue engineering. Additionally, it has been reported that solubilized elastin can not only induce angiogenesis, but also increase elastic fiber synthesis. Fibronectin Fibronectin is an adhesive glycoprotein that is primarily involved in cell– adhesive interactions. Intracellular signaling induced by cell adhesion on fibronectin plays a critical role in cytoskeletal organization, cell cycle progression, growth and cell survival and differentiation.27,28 Fibronectin plays a mayor role in the wound healing process and is the candidate material for tissue engineering since it is involved in ECM assembly. It is a glycoprotein that has the ability to bind integrins, collagens, fibrins and glycosaminoglycans (GAGs). Integrins such as α5β1 and αvβ3 are the major receptors that direct fibronectin matrix assembly; furthermore, α5β1 and αvβ3 integrins are central to regulating downstream events, including cell survival and cell-cycle progression.29 The fibronectin matrix stimulates fibroblast in-growth both in vitro and in vivo.30 In the literature it is well described that fibrin and fibronectin cross-link together in the process of blood clotting, which in turn forms the main structural network to support the wound until collagen is deposited.31,32

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Glycosaminoglycans GAGs are negatively charged polysaccharides of differing degrees of complexity that mediate many biological functions. GAGs are long, linear and highly negatively charged polysaccharides and consist of a backbone of repeating disaccharide units generally formed by a uronic acid and a glycosamine residue. Based on backbone structure, four major classes of GAGs are formed: heparan sulfate (HS)/heparin; chondroitin sulfate (CS)/dermatan sulfate (DS); keratan sulfate (KS); and hyaluronic acid (HA). The many biological characteristics of GAGs also make them valuable molecules for incorporation into matrices. For instance, HS is able to bind and modulate bioactive components like growth factors, cytokines and proteases, which are essential for basic cellular phenomena like cell adhesion, growth differentiation and activation, and implicate a role for GAGs in wound healing, inflammation, tissue morphogenesis and homeostasis. The incorporation of GAGs into a biomatrix promotes angiogenesis, tissue regeneration, reduces foreign body reactions and preserves matrix integrity in vivo (e.g. water absorption, osmotic swelling, tissue stiffness and strength).33 These sugar molecules with and without growth factors are crucial for the creation of micro- and nanoenvironments in the biomatrix for controllable cell growth.

Growth factors and cytokines Since the ECM serves as a reservoir for growth factors, and promotes their long-term bioavailability, smart scaffolds that replicate the ECM may be fortified with growth factors that regulate many aspects of cell behavior – such as growth, differentiation, migration, proliferation and survival. Growth factors can specifically induce biological phenomena like angiogenesis (by vascular endothelial growth factor (VEGF)) and the proliferation of certain types of cells including urothelial cells (by endothelial growth factor (EGF)) and fibroblasts (by basic fibroblast growth factor (bFGF)). Growth factors are stabilized and protected from proteolytic degradation by their interactions with ECM components such as GAGs.34 These combinations are ideal to mimic the growth factor binding and/or sustained release in biomaterials. There is a particular interest in an array of different growth factors in urological tissue engineering. For tissue engineering of the bladder for instance, angiogenic factors like bFGF and VEGF have great potential, whereas EGF could be used primarily to stimulate urothelial cells. Finally, the transforming growth factor-β (TGF-β) family is an example of multifunctional cytokines that regulate cell growth, differentiation and ECM formation. In mammals, there are three isoforms – TGFβ1, TGF-β2 and TGF-β3 – and although they have up to 85% amino acid sequence homology, there are

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known differences in their potencies and biological activities in vivo. For the development of smart biomatrices it is particular interesting that TGF-β1 and TGF-β2 are implicated in cutaneous scarring, whereas TGF-β3 has an anti-scarring effect.35

11.5

Future trends

A strategy has been proposed to isolate the main constituents of the ECM and to use them directly after purification, with or without further modifications. As ECM plays an instructive role in cell activities, the hypothesis here is that such biomolecules would maintain the biological information and other physicochemical features, which would preserve a potential space for new tissue development and cell in-growth. This would help to overcome one of the main drawbacks in the use of synthetic materials, which lack cell recognition signals. Since scaffolds are to be used inside the human body, special demands are put upon preparations used for this purpose. These include: purity (to avoid immunological response due to contaminants), porosity (for cellular in-growth), biocompatibility and biodegradability. Tissue engineering scaffolds made from conventional fabrication techniques are inadequate for the growth of thick cross-sections of tissue due to the diffusion constraints posed by foam structures; the three-dimensional architecture of the scaffold is of great importance for the survival and behavior of cells. In addition, controlling the purity and concentrations of these matrix molecules allows the design of a specific scaffold (or different combinations) on the drawingboard, followed by its construction in the laboratory. The scaffold architecture, for example its porosity, influences the mechanical properties of the scaffold; porosity can be increased by chemical, enzymatical or physical cross-linking. One method to control porosity is by freezing and lyophilizing techniques, whereby highly porous threedimensional bioscaffolds can easily be obtained.36,37 In general, combining the different components and techniques to prepare a construct allows for maximal flexibility and further components can be omitted and added at wish. For instance, in order to engineer urinary bladder, the top element of the bi-layered scaffold construct is made of non-porous collagen film (airdried) which is substituted with HS and HS-binding growth factor epidermal growth factor (EGF), to stimulate the proliferation of the urothelial cells (Fig. 11.1). On the other hand, the second layer, mimicking the matrix of the muscle layer, is composed of a porous collagen scaffold containing elastin fibers for elasticity, DS (the major GAG in bladder) and HS, with growth factors bFGF (proliferation of fibroblasts and stimulation of angiogenesis) and VEGF (stimulation of angiogenesis). It is anticipated that when such a scaffold is implanted at the site of the bladder or urinary tracts,

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Type I collagen film Elastin porous scaffold Type I collagen Heparan sulfate Dermatan sulfate VEGF bFGF EGF

11.1 Example of a ‘smart bladder construct’ consisting of non-porous collagen film with HS and EGF and a porous collagen–elastin matrix (50 : 50) with DS and HS combined with bFGF and VEGF.

cells will infiltrate and start producing the desired tissue. After a period of time, the scaffold will be degraded and replaced by new urinary tissue.

11.6

References

1 howe a, aplin ae, alahari sk and juliano rl: Integrin signaling and cell growth control. Curr Opin Cell Biol. 10: 220–31, 1998. 2 schwartz ma and baron v: Interactions between mitogenic stimuli, or, a thousand and one connections. Curr Opin Cell Biol. 11: 197–202, 1999. 3 streuli c: Extracellular matrix remodelling and cellular differentiation. Curr Opin Cell Biol. 11: 634–40, 1999. 4 mann bk, gobin as, tsai at, schmedlen rh and west jl: Smooth muscle cell growth in photopolymerized hydrogels with cell adhesive and proteolytically degradable domains: synthetic ECM analogs for tissue engineering. Biomaterials 22: 3045–51, 2001. 5 hubbell ja: Materials as morphogenetic guides in tissue engineering. Curr Opin Biotechnol. 14: 551–8, 2003. 6 halstenberg s, panitch a, rizzi s, hall h and hubbell ja: Biologically engineered protein-graft-poly(ethylene glycol) hydrogels: a cell adhesive and plasmin-degradable biosynthetic material for tissue repair. Biomacromolecules 3: 710–23, 2002. 7 murugan r and ramakrishna s: Design strategies of tissue engineering scaffolds with controlled fiber orientation. Tissue Eng. 13: 1845–66, 2007. 8 ellis dl and yannas iv: Recent advances in tissue synthesis in vivo by use of collagen-glycosaminoglycan copolymers. Biomaterials 17: 291–9, 1996.

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9 kim bs, baez ce and atala a: Biomaterials for tissue engineering. World J Urol. 18: 2–9, 2000. 10 daamen wf, van moerkerk ht, hafmans t, buttafoco l, poot aa, veerkamp jh and van kuppevelt th: Preparation and evaluation of molecularly-defined collagen-elastin-glycosaminoglycan scaffolds for tissue engineering. Biomaterials 24: 4001–9, 2003. 11 rosso f, marino g, giordano a, barbarisi m, parmeggiani d and barbarisi a: Smart materials as scaffolds for tissue engineering. J Cell Physiol. 203: 465–70, 2005. 12 furth me, atala a and van dyke me: Smart biomaterials design for tissue engineering and regenerative medicine. Biomaterials 28: 5068–73, 2007. 13 brown ra, blunn gw and ejim os: Preparation of orientated fibrous mats from fibronectin: composition and stability. Biomaterials 15: 457–64, 1994. 14 toba y, ajiki k, horie m, sango k and kawano h: Immunohistochemical localization of calbindin D-28k in the migratory pathway from the rat olfactory placode. J Neuroendocrinol. 13: 683–94, 2001. 15 nillesen st, geutjes pj, wismans r, schalkwijk j, daamen wf and van kuppevelt th: Increased angiogenesis and blood vessel maturation in acellular collagenheparin scaffolds containing both FGF2 and VEGF. Biomaterials 28: 1123–31, 2007. 16 kanematsu a, yamamoto s, noguchi t, ozeki m, tabata y and ogawa o: Bladder regeneration by bladder acellular matrix combined with sustained release of exogenous growth factor. J Urol. 170: 1633–8, 2003. 17 guan j, stankus jj and wagner wr: Biodegradable elastomeric scaffolds with basic fibroblast growth factor release. J Control Release 120: 70–8, 2007. 18 nuininga je, van moerkerk h, hanssen a, hulsbergen ca, oosterwijkwakka j, oosterwijk e, de gier rp, schalken ja, van kuppevelt th and feitz wf: A rabbit model to tissue engineer the bladder. Biomaterials 25: 1657–61, 2004. 19 el-kassaby aw, retik ab, yoo jj and atala a: Urethral stricture repair with an off-the-shelf collagen matrix. J Urol. 169: 170–3; discussion 173, 2003. 20 koh cj and atala a: Recent advances in the field of urology. Curr Urol Rep. 7: 43–9, 2006. 21 foran jp, patel d, brookes j and wainwright rj: Early mobilisation after percutaneous cardiac catheterisation using collagen plug (VasoSeal) haemostasis. Br Heart J. 69: 424–9, 1993. 22 van wachem pb, van luyn mj, olde damink lh, dijkstra pj, feijen j and nieuwenhuis p: Biocompatibility and tissue regenerating capacity of crosslinked dermal sheep collagen. J Biomed Mater Res. 28: 353–63, 1994. 23 leonard mp, decter a, mix lw, johnson hw and coleman gu: Endoscopic treatment of vesicoureteral reflux with collagen: preliminary report and cost analysis. J Urol. 155: 1716–20, 1996. 24 leonard mp, decter a, mix lw, johnson hw and coleman gu: Treatment of urinary incontinence in children by endoscopically directed bladder neck injection of collagen. J Urol. 156: 637–40; discussion 640–1, 1996. 25 okano t, satoh s, oka t and matsuda t: Tissue engineering of skeletal muscle. Highly dense, highly oriented hybrid muscular tissues biomimicking native tissues. ASAIO J. 43: M749–53, 1997.

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26 compton cc, butler ce, yannas iv, warland g and orgill dp: Organized skin structure is regenerated in vivo from collagen-GAG matrices seeded with autologous keratinocytes. J Invest Dermatol. 110: 908–16, 1998. 27 hynes ro: Cell-matrix adhesion in vascular development. J Thromb Haemost. 5 Suppl. 1: 32–40, 2007. 28 frisch sm and ruoslahti e: Integrins and anoikis. Curr Opin Cell Biol. 9: 701–6, 1997. 29 wu c, hughes pe, ginsberg mh and mcdonald ja: Identification of a new biological function for the integrin alpha v beta 3: initiation of fibronectin matrix assembly. Cell Adhes Commun. 4: 149–58, 1996. 30 currie lj, sharpe jr and martin r: The use of fibrin glue in skin grafts and tissueengineered skin replacements: a review. Plast Reconstr Surg. 108: 1713–26, 2001. 31 clark ra, dellapelle p, manseau e, lanigan jm, dvorak hf and colvin rb: Blood vessel fibronectin increases in conjunction with endothelial cell proliferation and capillary ingrowth during wound healing. J Invest Dermatol. 79: 269–76, 1982. 32 clark ra, quinn jh, winn hj, lanigan jm, dellepella p and colvin rb: Fibronectin is produced by blood vessels in response to injury. J Exp Med. 156: 646–51, 1982. 33 pieper js, van wachem pb, van luyn mja, brouwer la, hafmans t, veerkamp jh and van kuppevelt th: Attachment of glycosaminoglycans to collagenous matrices modulates the tissue response in rats. Biomaterials 21: 1689–99, 2000. 34 cartwright l, farhat wa, sherman c, chen j, babyn p, yeger h and cheng hl: Dynamic contrast-enhanced MRI to quantify VEGF-enhanced tissueengineered bladder graft neovascularization: pilot study. J Biomed Mater Res A. 77: 390–5, 2006. 35 shah m, foreman dm and ferguson mw: Neutralisation of TGF-beta 1 and TGFbeta 2 or exogenous addition of TGF-beta 3 to cutaneous rat wounds reduces scarring. J Cell Sci. 108 (Pt 3): 985–1002, 1995. 36 faraj ka, van kuppevelt th and daamen wf: Construction of collagen scaffolds that mimic the three-dimensional architecture of specific tissues. Tissue Eng. 13: 2387–94, 2007. 37 geutjes pj, daamen wf, buma p, feitz wf, faraj ka and van kuppevelt th: From molecules to matrix: construction and evaluation of molecularly defined bioscaffolds. Adv Exp Med Biol. 585: 279–95, 2006.

12 Natural biomaterials for urological tissue engineering C. C. R O T H and B. P. K R O P P, The University of Oklahoma Health Sciences Center, USA; and E. Y. C H E N G, Children’s Memorial Hospital in Chicago, USA

Abstract: Many tissue engineering protocols rely on an extracellular matrix or scaffold to support regenerating tissues. These biomaterials may be synthetic or naturally occurring. Biomaterials derived from natural sources include elemental alginate and collagen. Additionally, extracellular matrices derived from various organs and animal sources may be used to provide a biologically intact scaffold for use in regenerative protocols. This chapter highlights the fundamental and pre-clinical research that has been conducted using natural biomaterials for tissue regeneration. Natural biomaterials have demonstrated considerable promise and are likely to be an important element of future successful tissue engineering protocols. Key words: tissue engineering, urology, collagen, alginate, small intestinal submucosa, bladder extracellular matrix.

12.1

Introduction

Tissue engineering combines medical technology with the native regenerative properties of living organisms. Through well-integrated processes, tissue engineering allows for greater regenerative potential than can be found in nature alone. Biomaterials are an integral component of tissue engineering. In select circumstances biomaterials may be sufficient to replace native tissues obviating the need for complex tissue regeneration. Alternatively, biomaterials may be implemented in a tissue engineering protocol in which the biomaterial serves as a temporary substitute in form or function. In this manner, the biomaterial may act as a tissue scaffold and maintain homeostasis while the complex process of tissue regeneration takes place. Regarding regeneration of whole tissues or organs, the ideal biomaterial will not promote an immunogenic response, will promote a regenerative process, and will be either reabsorbed by or integrated into the target tissue. Both synthetic and natural biomaterials, which meet these criteria, have been used. 255

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The purpose of this chapter is to introduce the natural biomaterials that are commonly used in urological tissue engineering. The historic use of natural biomaterials will be outlined. We will discuss the source and fundamental properties of collagen and alginate as well as the processing of these biomaterials into gels and matrices. We will also highlight the production and basic properties of the various naturally derived collagen extracellular matrices. Pre-clinical applications of these natural biomaterials will be covered in this chapter with organ-specific tissue engineering protocols to be addressed elsewhere in the text.

12.2

Historical application of natural biomaterials

The history of urological surgery is replete with cases of bladder substitution with various forms of xenogeneic, allogeneic, and autogeneic biomaterials. Seromuscular grafts (Shoemaker, 1955), omentum (Goldstein and Dearden, 1966), peritoneum (Jelly, 1970), dura (Kelâmi, 1971), placenta (Fishman et al., 1987), and pericardium (Kambic et al., 1992) have all been employed as bladder wall substitutes. In theory, partial bladder substitution would simply require an impermeable membrane (synthetic or natural) to replace the deficient portion of the bladder. The resulting increase in bladder capacity could allow for passive storage of urine with the potential for maintained continence. The concept of regenerating a functional bladder was not emphasized by earlier researchers and the results were inconsistent. As research with biomaterials progressed, the importance of tissue regeneration became evident. Lyophilized human dura was implemented by Kelâmi (1971) as a bladder wall substitute in an experimental dog model and subsequently in seven patients following partial cystectomy. In both groups, initial results demonstrated adequate capacity and function. Dura patches have been used sparingly since the initial description with one subsequent patient developing transitional cell carcinoma over the dura portion of the bladder 30 months after implantation (Selli et al., 1986). In addition to bladder substitution, dura has been used for replacement of tunica albuginea in Peyronie’s disease and for urethral suspension for stress incontinence (Kelâmi, 1976). All of these applications rely on the mechanical properties of dura with no dependence on the regenerative properties of the graft. As research progressed, investigators recognized the importance of regeneration of native bladder with subsequent degradation of the implant material. Kambic et al. (1992) investigated the use of acetylated bovine pericardium as a bladder patch in a canine model with the theory that the pericardium would be replaced by regenerated bladder. Dogs were followed for up to 3 years post-implantation and demonstrated stable or increased bladder capacities. While re-epithelialization was noted on the

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luminal surface of the graft, scar formation at the apex of the graft and failure of organized smooth muscle development were also observed. The use of placental membranes as a biodegradable bladder substitute has also been investigated (Fishman et al., 1987). Canine bladders augmented with human placental membranes demonstrated improved capacity over control animals and also exhibited evidence of urothelial and destrusor regeneration. The placental membranes could not be identified in later surgical specimens. Despite initially promising results, follow-up data are lacking on all of the above natural grafts. It is possible that long-term studies were either not conducted or demonstrated poor outcomes and were, therefore, never published.

12.3

Fundamental biomaterials

12.3.1 Alginate Augst et al. (2006) have outlined the processing of alginate gels and reviewed the current applications of alginate hydrogels in tissue engineering. Alginates are polysaccharides isolated from common coastal algae such as Laminaria hyperborea and Laminaria lessonia. In 1881, Stanford first isolated alginate by extracting it with sodium carbonate and then precipitating it out of solution at low pH. Alginate was later characterized to be a block polymer composed of d-mannuronic acid and guluronic acid (Fig. 12.1). The ratio of d-mannuronic acid and guluronic acid will vary by algae source and by precipitation method. The polymers can be processed into gels by a method of oxidization and covalent or ionic cross-linking. Processing the gels under various conditions can alter the chemical properties of the gels to match the intended application. When used as biomaterial, alginate gel offers advantages over alternative materials. The hydrophilic nature of alginate gels makes them biocompatible and non-immunogenic. Additionally, alginate possesses a gentle gelling

–OOC

OH

O OH

O

G

O OH

OH OH O

O

HO

–OOC

O OH

–OOC O

O

O

OH –OOC G

M

M

12.1 Alginate polysaccharide that consists of two guluronic (G) and two mannuronic (M) acid residues with (1,4)-linkages. Reprinted from Drury et al. (2004), with permission from Elsevier.

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behavior that allows for encapsulation of transplanted cells and pharmaceuticals with minimal trauma. Alginate gel, while composed of natural materials, is not a natural component of higher biological systems. This unique feature introduces potential disadvantages including the inability of most cells to naturally bind to the gel and the lack of regulated enzymatic degradation. Cell binding can be improved by the addition of cellular adhesion molecules such as arginine – glycine – aspartic acid, RGD, to the gel. Altering the amount of cross-linking will influence the rate of degradation (Augst et al., 2006). Within the field of urology, alginate hydrogels have shown promise as a bulking agent (Thornton et al., 2004) and as delivery vehicles for both cultured cells (Machluf et al., 2003) and pharmaceuticals (Oztürk et al., 2004). Various bulking agents such as collagen and dextranomer/hyaluronic acid are currently being utilized for the management of incontinence and vesicoureteral reflux (VUR). The utility of alginate hydrogels as a bulking agent has been investigated. Animal studies have shown that alginate gel can be dehydrated, compressed, and injected into subepithelial locations. Once rehydrated with saline, the gels obtain their original volume and shape. The resulting volume is maintained for up to 6 months (Thornton et al., 2004). The actual use of alginate gel in the treatment of VUR has not been demonstrated. However, alginate solution has been utilized as a scaffold for cultured chondrocytes in the endoscopic treatment of VUR. Atala et al. (1994) demonstrated that an alginate/chondrocyte solution when injected in a subureteral location can correct VUR in mini-pigs. Later studies in humans demonstrated that VUR was corrected in 83% of children with grade II to IV VUR following up to three injections of a cultured chondrocyte and alginate solution (Diamond and Caldamone, 1999). One year following treatment, 65% of patients remained free of reflux (Caldamone and Diamond, 2001). Owing to their inert nature and lack of enzymatic degradation, alginate hydrogels can be utilized as protective carriers for the delivery of cultured cells and pharmaceuticals. The use of alginate hydrogel as a carrier for cultured Leydig cells was investigated in a rat model. Castrated rats were administered subcutaneous injections of Leydig cells or Leydig cells in an alginate carrier. Those animals treated with Leydig cells alone demonstrated no increase in testosterone levels compared with the animals treated with alginate-encapsulated Leydig cells which achieved a serum testosterone level of 40% of normal. The increase in testosterone was evident for over 40 days (Machluf et al., 2003). Alginate gels have been proposed as carriers for time-released pharmaceuticals. An advantage of using carriers is that drugs can be delivered to site-specific locations and the dose can be maintained at physiological levels for a longer duration. One proposed scenario is the administration of intra-

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vesical chemotherapy following resection of bladder tumors. The drug carrier could bind to the urothelium and release lower doses of the drug over time in an effort to minimize side effects and maximize efficacy. Toward this end, Oztürk et al. (2004) have preliminarily investigated the use of alginate gels as a carrier for Mitomycin-C. While the technical ability to combine the drug with alginate was demonstrated, no therapeutic outcomes are available.

12.3.2 Collagen Collagen is the most abundant and ubiquitous protein in the human body. It is composed of three polypeptides (α chains), each with an amino acid sequence of (-glycine-X-Y-)n, where X is commonly proline and Y is frequently hydroxyproline. Within the extracellular space collagen does not exist as isolated molecules. Rather, the molecules aggregate into fibrils, which are the building units for large collagen fibers (Fig. 12.2). The amino acid composition of the polypeptides can vary to result in 19 distinct collagen proteins. Type I collagen is the most abundant form and is a major constituent of bone, skin, ligament, and tendon. The three-dimensional organization of the collagen fibers will vary by anatomic location. Arrangement of collagen into parallel fibers, as seen in tendons, provides high tensile strength. Fibers in skin are arranged in random arrays that provide resiliency to tissues under stress (Li, 2000). In addition to the abundance in which collagen is found, several other characteristics of collagen make it attractive for tissue engineering. Collagen is a common component of the extracellular matrix (ECM) in all tissues and organs, and many cells bind specifically to collagen. Utilizing this property of cell–collagen interaction is beneficial for tissue engineering. Platelets will also bind to exposed collagen, which imparts an intrinsic hemostatic property to elemental collagen. Additionally, the most important property of collagen is that it is a poor immunogen. There are few interspecies differences in the structure of type I collagen, and this homology allows for collagen to be used as a xenogeneic graft without significant development of an immune response within the host (Li, 2000). A soluble collagen solution can be prepared from a collagenous tissue by using proteolytic enzymes like pepsin. Pepsin cleaves telopeptides, the natural cross-linking sites of collagen, to produce small soluble collagen aggregates that are purified by precipitation with neutral salts. Once in the soluble form, collagen can be reorganized into a variety of matrices including membranes, sponges, and gels (Li, 2000). Collagen gels were initially investigated for urological use as a barrier to prevent urine extravasation following ureterotomy. Gorham et al. (1984) developed a collagen film by processing bovine hide and demonstrated that

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Microfibrils

Hole zone Packing of molecules

Collagen molecule

280 nm

α2 Triple α1 helix α1

1.5 nm

Collagen α-chain (purity to molecules)

12.2 Microstructure of collagen showing the three α-chains of the triple helix. Three separate α-chains are wrapped around each other to form a rope-like triple-stranded, helical rod. Every third amino acid of the native molecule is glycine. This figure was published in DeLee and Drez’s Orthopaedic Sports Medicine, (Best et al., 2003) and is reprinted with permission from Elsevier; Copyright Elsevier (2003).

the film would not calcify when exposed to urine in vitro. The collagen film was then used to cover ureterotomies in rabbits and the experimental group demonstrated decreased evidence of urinary extravasation (Scott et al., 1986). This study demonstrated two important properties of collagen gels: the gels were rapidly biodegradable and the gels were difficult to work with surgically owing to their fragile nature. In order to overcome the latter problem, Monsour et al. (1987) combined polyglactin mesh with the collagen gel to create a more durable graft. These grafts were subsequently used to repair bladder defects with histological evidence of urothelium and smooth muscle regeneration in a background of scar formation.

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As tissue engineering techniques evolved, interest developed in using collagen gels as a carrier for cultured urothelial and smooth muscle cells. Chlapowski et al. (1983) demonstrated the ability to grow urothelial cells on a collagen substrate. Considering the limitations of using collagen gels alone, Hakim et al. (1994) explored the utility of culturing urothelial cells on composite grafts of collagen and polyglactin. The authors demonstrated poor growth on polyglactin grafts alone and 90–100% confluence on composite collagen–polyglactin grafts. This example illustrates the inherent nature of collagen to promote cell growth. Researchers have continued to pursue a clinical application for collagen gels. Advances have been made in both the manufacturing process and surgical implementation of collagen products. There are multiple commercial sources for collagen gels and sponges. Independent research has demonstrated that these products are compatible with cultured urothelial and bladder smooth muscle cells (Becker et al., 2006; Danielsson et al., 2006). Newer generation collagen gels have demonstrated better surgical handling. Nuininga et al. (2004) fabricated a cross-linked collagen gel that was then used to augment rabbit bladders. The collagen gel demonstrated durable physical properties with no evidence of urinary extravasation or premature breakdown. The collagen gel produced regeneration of the mucosa and muscularis. Encrustation of the collagen gel was noted in nearly half of the rabbits. The authors proposed adding glycosaminoglycans (GAGs) to the collagen matrix in an effort to limit encrustation. The ability to alter the composition of collagen matrices will allow continued investigation into the optimal biomaterial.

12.4

Collagen-based extracellular matrices

Collagen is the most abundant protein within the ECM; however, the ECM is a complex substrate containing many additional elements that are essential for tissue growth and differentiation. Similar to collagen, many of these components may be produced in a laboratory setting. As demonstrated with elemental collagen, tissue regeneration with any one component is likely to be challenging. For this reason, intense research into the use of naturally occurring ECM grafts as a scaffold for tissue regeneration is ongoing. Each ECM is unique to the tissue from which it is derived though the same basic components will be present regardless of origin. Common components include collagen, fibronectin, laminin, GAGs, and growth factors (Badylak, 2004). Type I collagen is the most prevalent of the collagen subtypes though each of the other 18 subtypes can be present based on the function of the tissue. Fibronectin is second only to collagen in quantity within the ECM. Fibronectin possesses ligands essential for cell binding and

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is important for normal biological development. Laminin is important in embryological development and is proposed to play a role in wound healing versus scar formation in adult tissues. GAGs, such as heparin sulfate and hyaluronic acid, are present in various quantities within the ECM depending on the tissue of origin. GAGs bind growth factors and promote retention of water, which gives the ECM a gel-like property. An extensive number of growth factors are found within the ECM. While the concentration of these factors can be adjusted experimentally, the naturally occurring ratio of growth factors is likely already optimized for tissue development and growth. Within the field of urology, several particular ECMs have been utilized for tissue regeneration. These include porcine small intestinal submucosa (SIS), bladder ECM, and dermal ECM. The use of bladder ECM to regenerate bladder is an example of using homologous grafts, either xenogeneic or allogeneic, to promote growth of the same organ in a different host. This theory has also been applied in using ureteral and vaginal ECM to regenerate ureters and vaginas, respectively.

12.4.1 Porcine small intestinal submucosa Fundamentals of small intestinal submucosa SIS was originally investigated for use as a vascular graft (Badylak et al., 1989). Based on the initial results, which demonstrated repopulation of the graft with endothelial cells, the application of SIS was expanded to other organ systems. Canine SIS was used to perform bladder augmentation in pigs and epithelialization of the graft was noted (Knapp et al., 1994). Since these initial studies, SIS has been used extensively throughout the genitourinary system. Badylak et al. (1989) described the technique of preparing SIS grafts. Following removal of mesenteric tissue, the bowel segment is everted and mechanical abrasion (scalpel handle, wet gauze, etc.) is used to remove the epithelium and lamina propria manually leaving the stratum compactum of the tunica mucosa intact. The bowel segment is then returned to the native orientation and the process is repeated to remove the tunica serosa and tunica muscularis externa. The remaining tissue (100 μm thick) will be a translucent white graft consisting of the original stratum compactum and muscularis mucosa of the tunica mucosa. The stratum compactum represents the luminal side of the graft. The graft is typically preserved by rinsing with saline and sterilizing with dilute (0.1%) peracetic acid. Alternative sterilization techniques include ethylene oxide, gamma irradiation, and E-beam irradiation (Badylak, 2004). Generally, no additional enzymatic treatment or chemical modification is employed.

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Porcine SIS is currently commercially available as Surgisis® (Cook Biotech Inc., West Lafayette, IN, USA). The exact manufacturing protocol is not available to the public though it is similar to the original method described above (M. Hiles, personal communication, 25 July 2007). The commercial product is subject to regulation as a medical device, and the source animals are raised under strict medical regiments and must pass United States Department of Agriculture inspection. The SIS is treated with peracetic acid and other solutions, which render it virus-free (Hodde and Hiles, 2001) and low in bacterial count and endotoxins. The SIS is dried by lyophilization and terminally sterilized with ethylene oxide. With proper storage, the strength, sterility, and biological activity are maintained for at least 2 years. As an ECM, SIS has been proven to contain the important elements needed for tissue growth including several types of collagen, GAGs, glycoproteins, and growth factors (Hodde et al., 2007a). In order to create a commercially available product that is safe for general use, a proven method of processing and sterilization must be employed. Hodde et al. (2007a) conducted a series of experiments to determine if the process of treating with peracetic acid, lyophilization, and sterilizing with ethylene oxide would significantly change the biochemical make-up of the SIS. They found that fibronectin and GAGs are retained following oxidation with peracetic acid and alkylation using ethylene oxide gas. Regarding the presence of growth factors, the amount of vascular endothelial growth factor (VEGF) decreases and the amount of fibroblast growth factor-2 (FGF2) is stable following treatment. Additionally, the three-dimensional architecture of the native SIS was not changed following processing. In a similar series of experiments the biological activity of the SIS was assessed (Hodde et al., 2007b). Following the processing procedure, fibroblasts were able to attach to the SIS, remain viable, and upregulate their expression of VEGF. Based on these observations, it appears that the commercial processing method for SIS does not significantly change the inherent properties of the ECM. The function of the small intestine varies by location. Just as the jejunum functions differently than the ileum, the fine structure of the ECM will also vary by location. Raghavan et al. (2005) conducted a series of in vitro studies to characterize physical differences in hand-made SIS obtained from proximal small bowel, hand-made SIS obtained from distal small bowel, and commercially available SIS (Surgisis®). Analysis of the microarchitecture by electron microscopy (Fig. 12.3) demonstrated that the commercially available SIS possessed a fibrous appearance of the serosal surface. This differed from the mucosal surface of the three grafts and the serosal surface of the hand-made SIS grafts, which appeared considerably less porous. Permeability studies revealed that the distal SIS demonstrated the least permeability to urea and that the mucosal surface of each of the grafts was

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12.3 Scanning electron photomicrographs demonstrating variations in physical structure of SIS depending on source and location of graft harvest. (a) Mucosal surface of COOK® SIS; (b) mucosal surface of hand-made SIS derived from proximal small bowel; (c) mucosal surface of hand-made SIS derived from distal small bowel; (d) serosal surface of COOK® SIS; (e) serosal surface of hand-made SIS derived from proximal small bowel; (f) serosal surface of handmade SIS derived from distal small bowel.

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less permeable than the serosal surface. These findings confirmed that the physical properties of SIS are location dependent. Porcine SIS has been used in a number of different organ systems across several species. Common to all of the experiments conducted with SIS is that graft rejection is not encountered. One likely explanation is that the components of the ECM are preserved across species. In order to better characterize the immune response to SIS, Allman et al. (2001) compared the immune response of various grafts implanted into the flank of mouse. The grafts used included porcine SIS, urinary bladder submucosa, rat muscle (xenogeneic control), and mouse muscle (syngeneic control). Histological analysis of the xenogeneic implants demonstrated intense inflammation consistent with rejection. The syngeneic graft and SIS implant demonstrated only mild inflammation. The cytokine and immunoglobin profile for the SIS-implanted mice demonstrated evidence of a Th2 immune response, which is typically associated with transplant graft acceptance. A complete analysis for bladder submucosa was not published, although the antibody profile was similar to that seen for SIS. These findings correlate with the clinical observation that xenogeneic SIS grafts do not undergo rejection by the host. A proposed advantage of ECM grafts is the ability of cells to adhere naturally to the three-dimensional structure. As noted in the preceding discussion, fibroblasts can be cultured on processed SIS. Additional studies have demonstrated the ability to culture urothelial cells and bladder smooth muscle cells on SIS. Zhang et al. (2000) co-cultured urothelial cells on the mucosal surface and bladder smooth muscle cells on the serosal surface of commercially available SIS (Surgisis®, Cook Biotech Inc., West Lafayette, IN, USA). Histological analysis of the co-cultured grafts demonstrated three-dimensional organization of the mucosa and smooth muscle with penetration of the cells into the center of the graft. This differed from the less-organized appearance of the cultured cells when only one cell type was grown on the graft or the cells were grown together on the same side of the graft. Despite this observation other investigators have had limited success with culturing urothelial cells on SIS. Feil et al. (2006) also used commercially available SIS (Surgisis®ES SLH-4S, Cook Deutschland, Mönchengladbach, Germany), but were not successful in culturing urothelium. Whether the difficulty in culturing the cells was due to the substrate on which the cells were grown or the fastidious nature of the cells is difficult to determine. Pre-clinical studies Following the initial work done by Knapp et al. (1994), SIS bladder augmentation was carried out in several animal models. Kropp et al.

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(1995) performed bladder augmentation in 22 rats. The experimental protocol consisted of excising the bladder dome and augmenting the bladder with a 1-cm2 portion of SIS. The bladders were harvested at time points ranging from 2 weeks to 11 months to provide a temporal outline of the regenerative process. Histological examination revealed that the mucosa would completely cover the graft by 2 weeks post-implantation. The remaining components of the bladder subsequently regenerated so that by 11 months post-implantation the regenerated portions of the bladder were indistinguishable from the native bladder. In vitro contractility studies and staining for nerve regeneration demonstrated that the SISregenerated bladder had normal contractility and nerve regeneration (Vaught et al., 1996). Based on promising results in the rat model, similar experiments were conducted in a dog model. Complete regeneration was noted though the smooth muscle in the regenerated portion of the bladder was less organized than the native smooth muscle. Cystometrograms demonstrated a stable bladder capacity, compliance, and post-void residual. Gross inspection of the bladders revealed no evidence of graft shrinkage, scarring, or development of bladder calculi (Kropp et al., 1996a). Further in vitro testing confirmed that the contractility of the regenerated bladder was similar to that of the native bladder. Additionally, muscarinic, adrenergic, and purigenic receptors were expressed in a normal pattern (Kropp et al., 1996b). As more experience with SIS bladder augmentation was developed, it became evident that tissue regeneration was not consistent across experiments. Osseous metaplasia was demonstrated in a majority of one series of dogs undergoing SIS bladder augmentation (Pope et al., 1997). Continued observation led researchers to conclude that four-ply SIS, electron beam sterilization, and utilizing proximal segments of bowel for SIS preparation all led to poor tissue regeneration (Kropp et al., 2004). Poor regeneration was also demonstrated in a model of severe bladder disease. In this particular experiment, 90% of the canine bladder was resected and the bladder was closed. One month later the bladders were augmented with either cellseeded or unseeded SIS. Evaluation 5–9 months later demonstrated incomplete regeneration and graft shrinkage. The authors concluded that the vascularity of a severely damaged bladder may not be able to sustain largescale tissue regeneration (Zhang et al., 2006).The use of SIS for bladder augmentation has demonstrated significant potential, although additional study is needed to produce reliable and predictable results. In addition to bladder surgery, SIS has been used as a graft for reconstruction of the urethra and tunica albuginea. Kropp et al. (1998) performed SIS grafting of the urethra in male rabbits. The SIS was used as an onlay graft and demonstrated regeneration of the urethral mucosa on a background of

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fibrous tissue with partial development of smooth muscle. No strictures or fistulae were demonstrated on post-operative urethrograms. SIS has been used as a patch graft for defects of the tunica albuginea. Weigel et al. (1996) performed implantation of SIS following elliptical excision of native tunica in 20 rats. Gross and histological examination 1–24 weeks later demonstrated no graft shrinkage and neovascularization of the graft with incorporation into the native corporal tissue. Clinically, SIS has now been used as both a corporal graft (Kropp et al., 2002) and urethral graft (Fiala et al., 2007). Ureteral replacement with SIS has been attempted in numerous preclinical trials with disappointing results to date. Jaffe et al. (2001) utilized SIS as a graft to replace an 11-mm segment of excised ureter. Epithelialization of the SIS was noted, though on a background of fibrous tissue with partial regeneration of smooth muscle. The ureters remained patent at the time of euthanasia (35 days). Later attempts to use SIS to replace a 4-cm segment of excised ureter in dogs resulted in partial tissue regeneration, although all ureters were completely occluded by 12 weeks (El-Assmy et al., 2004). Better results were achieved when SIS was used as an onlay graft to repair partially resected ureters in pigs. Nine weeks following surgery, the ureters remained patent without evidence of stricture or hydronephrosis (Smith et al., 2002). The hypothesis that cell-seeded SIS would result in better tissue regeneration and less scarring was evaluated by El-Hakim et al. (2005). In one arm of the study a 5-cm segment of ureter was excised in pigs and replaced with SIS. An additional arm consisted of replacing a 3-cm segment of ureter in dogs with SIS seeded with bladder epithelial cells. Both groups developed complete stenosis of the ureter despite histological evidence of tissue regeneration. Continued investigation is needed before SIS can be considered as a viable option for ureteral replacement. SIS has been used in surgical treatment of stress urinary incontinence and pelvic organ prolapse. In this setting the SIS is used for its structural support as much as its regenerative potential. Cannon et al. (2005) have evaluated the use of cell-seeded SIS as an additional treatment for urinary incontinence. They used SIS seeded with muscle-derived cells to create a suburethral sling in the rat model. They demonstrated that the technique was surgically feasible and increased the leak point pressure in treated rats. Histological examination demonstrated that muscle-derived cells were incorporated into the graft. The regenerative properties of injectable SIS have also been evaluated for potential treatment of stress incontinence and VUR. Experimentation in dogs demonstrated that submucosal injection of SIS resulted in de novo smooth muscle formation (Furness et al., 1999). Refinement of these techniques may eventually result in a more durable treatment option for both incontinence and VUR.

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12.4.2 Bladder extracellular matrix Fundamentals of bladder extracellular matrix Following the initial results achieved with SIS for bladder augmentation, Probst et al. (1997) hypothesized that homologous ECM may provide even more complete regeneration than seen with SIS. The authors used rat bladder ECM termed bladder acellular matrix graft (BAMG), to perform bladder augmentations in rats and found that by 12 weeks post-implantation the regenerated bladder was nearly indistinguishable from native bladder. Based on the success of the initial experiment, researchers have continued to evaluate BAMG as a graft for bladder regeneration. Additionally, homologous grafts have been produced in a similar fashion for other organs including the ureter and vagina. The preparation of BAMG was outlined by Probst et al. (1997). Briefly, the bladder dome was excised and washed with phosphate-buffered saline (PBS) and 0.1% sodium azide. The mucosa was removed by mechanical abrasion. Partial cell lysis was achieved with incubation in PBS and 0.1% sodium azide for 5–6 hours. Cell lysis was completed by incubation in sodium chloride and 2000 Kunitz units of DNase (Sigma, St Louis, MO, USA) for 6–8 hours. The processing was completed by washing the samples with 4% sodium desoxycholate containing 0.1% sodium azide to stabilize the lipid membrane. The grafts were stored in 10% neomycin sulfate until the time of use. This process produces an acellular ECM. As an alternative technique, BAMG and other homologous grafts may be produced without mechanical abrasion of the mucosa. Utilization of the full chemical process alone can result in an acellular graft (Sievert and Tanagho, 2000). Different institutions employ slight variations in the manufacturing process (Merguerian et al., 2000) without any appreciable difference in graft quality. The effects of graft preparation on inherent growth factors and scaffold structure have not been reported. Bladder ECM is commercially produced by Acell Inc. (Columbia, MD, USA). Urinary bladder matrix (UBM) is marketed for both human and veterinary applications. Per the product information, use of UBM is not limited to the genitourinary system. Acell Inc. has received Food and Drug Administration (FDA) clearance for the human applications of their ECM products (acell.com, 2007). The manufacturing process of UBM is not available to the public. Bladder ECM differs from SIS in that research has been conducted on bladder ECM derived from multiple species. While the ultastructure of bladder ECM is similar across species, slight variations in composition have been noted. Electron microscopy demonstrated that type I and III collagen, as well as elastin, are the major components of bladder ECM. The

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ratio of type I to III collagen varies by species with rat bladder ECM containing higher levels of type I collagen than human or pig bladder ECM. Pig and human tissue demonstrated higher levels of type III collagen and elastin (Dahms et al., 1998a). Later evaluation demonstrated slight differences in collagen I:III ratios among monkeys, dogs, and humans (Sievert et al., 2006). Dahms et al. (1998a) performed biomechanical testing on bladder ECM from rats, pigs, and humans. They demonstrated no significant difference in stress, strain, or elastic modulus between bladder ECM and native bladder in pig and human tissue. Rat bladder ECM demonstrated a significant decrease in tensile strain and elastic modulus when compared with native rat bladder. This finding may be explained by the loss of smooth muscle and mucosa. While not statistically significant, the ultimate tensile strengths of rat bladder and rat bladder ECM were higher than the same measures in the other species. This finding could be attributed to a higher level of type I collagen, which is generally less elastic than type III collagen. Some properties of bladder ECM are not as well studied as the properties of SIS. No information is available regarding the inherent growth factors within bladder ECM. Likewise, the immune response to bladder ECM was only partially studied by Allman et al. (2001) in their evaluation of the immune response to SIS. As previously discussed, xenogeneic SIS grafts, like syngeneic grafts, fail to illicit a rejection response. The immune response to SIS is limited to a Th2 response, which is linked to transplant acceptance. The antibody expression demonstrated in a Th2 response is limited to immunoglobulin G1 (IgG1). In this evaluation, the antibody produced following exposure to bladder ECM or SIS was IgG1. The authors concluded that the same immune response should be observed following exposure to any ECM regardless of the source organ. The biocompatibility of bladder ECM and cultured cells has been evaluated in several studies. Pariente et al. (2001) demonstrated the effects of various bioscaffolds, natural and synthetic, on cultured urothelial cells. Bladder ECM proved not to be cytotoxic and allowed for growth of cultured urothelial cells directly on the scaffold. The same authors repeated the experiment with bladder smooth muscle cells and again demonstrated cell compatibility with bladder ECM (Pariente et al., 2002). Brown et al. (2005) conducted co-culture experiments with urothelial cells and bladder smooth muscle cells on bladder ECM. They were able to demonstrate that both cell lines grew well on the bladder ECM. When the cell lines were grown together on the scaffold, smooth muscle cells demonstrated increased infiltration into the graft and increased contractility compared with when smooth muscle cells were grown alone.

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Pre-clinical studies The use of bladder ECM was pioneered by the group of researchers led by Dr E. A. Tanagho at the University of California School of Medicine, San Francisco. Probst et al. (1997) outlined the preparation of BAMG and demonstrated that bladder augmentation with homologous rat BAMG resulted in regeneration of all of the components of the bladder wall without evidence of scarring. Intravesical stones were found in 53% of the animals. Using the same model of regeneration, Piechota et al. (1998a) evaluated the in vitro functional properties of the regenerated bladder. The BAMG regenerated bladder demonstrated a functional response to both electrical field stimulation and exposure to smooth muscle stimulants. The magnitude of muscle contraction was less than that seen with native bladder although it increased gradually as regeneration matured at later post-implantation time points. Histological analysis demonstrated qualitatively indistinct regeneration compared with native bladder. Cholinergic and adrenergic nerve fibers were identified as regeneration progressed. Again a high rate of stone formation (80%) was observed with the most common stone composition being struvite. An additional investigation by the same group sought to establish in vivo functionality of xenogeneic BAMG-augmented bladders. Hamster, rabbit, and dog BAMG were used to augment rat bladders. Histologically, all grafts resulted in regeneration without evidence of rejection. Partial detrusor regeneration was demonstrated in all BAMG grafts. When compared with a control partial cystectomy bladder, all regenerated bladders for each type of BAMG demonstrated increased functional capacity and improved compliance. These evaluations demonstrated the feasibility of BAMG augmentation and further established no significant difference in regeneration when xenogeneic grafts are employed (Piechota et al., 1998b). The molecular pathway of tissue regeneration is incompletely understood. Prior research has identified the transforming growth factors, TGF-α and TGF-β, as potential mediators of epithelial–mesenchymal interactions (Baskin et al., 1996). In an effort to evaluate growth factors that may be relevant to the regeneration pathway, Dahms et al. (1998b) used BAMG in a rat model to demonstrate the expression of TGF in regenerated bladders versus control partial cystectomy bladders. In the BAMG-regenerated tissue, mRNA levels of TGF-α and TGF-β were significantly elevated over control levels from 2 weeks to 6 months following graft implantation. While only a limited evaluation, this finding illustrates the potential importance of molecular pathways in tissue engineering models. Attempting to approximate the conditions in which BAMG may eventually be used clinically, Urakami et al. (2007) evaluated bladder regeneration in a spinal cord injury (SCI) rat model. The experimental rats underwent

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transection of the lower thoracic spinal cord 8 weeks prior to augmentation with BAMG. Urodynamics were performed immediately prior to augmentation and 8 weeks post-operatively. Spinal cord transection resulted in hyperreflexic bladders in 71% of the rats and underactive bladders in 29%. The average bladder capacity and post-void residual for the SCI rats was significantly increased when compared with capacities in normal rats. Following augmentation with BAMG the SCI rats demonstrated a significant decrease in uninhibited bladder contractions and post-void residual and improved compliance. Histologically, complete regeneration was achieved in the SCI rats with no difference observed when compared with historic controls. It is not possible to create a model that emulates the entire scope of neurological lesions, although this fundamental study provides insight into potential clinical applications for bladder ECM. Additional studies evaluated the role of bladder ECM grafts for bladder regeneration in a large animal model. Mongrel dogs underwent a 20–50% partial cystectomy with replacement with a BAMG graft that was 150% of the size of the resected bladder. Urodynamics were performed just prior to surgery and at the time of bladder harvest – 7 months post-operatively. Findings were compared with control dogs that underwent 40% partial cystectomy and bladder closure without grafting. Cystometric data revealed that control dogs had a statistically significant decrease in bladder capacity at 7 months post-operation compared with pre-operative values. The augmented dogs demonstrated a significant increase in bladder capacity. Gross examination demonstrated that the grafts were 70% of their original size and covered 105.2% of the resected area. No stone formation was noted. The histology of the regenerated bladders was similar to that found in previous studies with the exception that the center of the graft had a higher percentage of collagen and lower density of smooth muscle compared with the edges of the graft. Following the results of this study, the authors expressed hope for a clinical trial utilizing BAMG for bladder augmentation although to date no follow-up data are available (Probst et al., 2000). Using a different detergent and enzyme preparation, the researchers at The Hospital for Sick Children and the University of Toronto also produced bladder ECM that they termed bladder acellular matrix allograft (BAMA) (Merguerian et al., 2000). The initial study evaluated the large-segment substitution (mean graft size 43.88 cm2) of BAMA in pig bladders. All animals survived the surgery and gross evaluation at the time of bladder harvest (4 weeks post-operatively) demonstrated no evidence of diverticula or stones and 10% graft shrinkage. Regeneration was ongoing with complete formation of the urothelium and partial development of bladder smooth muscle. In a second study, which also used a graft over 40 cm2, the evaluation was carried out to 12 weeks. At the end of the study, graft size was decreased

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by 31% of the original size. Smooth muscle was still less organized than control tissue although neural regeneration was evident (Reddy et al., 2000). An additional investigation using large segments of BAMA was carried out until 22 weeks post-implantation. Graft shrinkage continued to progress and the center of the graft remained devoid of organized smooth muscle (Brown et al., 2002). This series of evaluations highlights some of the limitations of ECM grafting in large animal models. Yoo et al. (1998) evaluated the effect of cell seeding on the regenerative properties of bladder ECM. The authors’ lab used a microdissection as opposed to chemical treatment to produce bladder ECM from dog bladders, which they termed allogenic bladder submucosa. Partial cystectomies were then performed in beagle dogs with half of the animals receiving a graft seeded with urothelial cells and bladder smooth muscle cells and the remaining animals receiving only allogenic bladder submucosa. Urodynamic evaluation up to 3 months following augmentation demonstrated a 99% increase in bladder capacity for bladders augmented with cell-seeded grafts compared with a 30% increase in capacity for bladders augmented with grafts without cultured cells. While the histological evaluation at 2–3 months post-operation noted that both grafts resulted in equivalent tissue regeneration, the authors attributed the higher capacity in the cellseeded group to prevention of graft shrinkage by the cultured cells. The authors have continued to utilize bladder ECM in their tissue engineering investigations including the only completed clinical trial for bladder regeneration. Attempts to regenerate urethral tissues have utilized bladder ECM grafts as well as urethral ECM grafts. Sievert et al. (2000, 2001) conducted a series of experiments where regeneration of the rabbit urethra was attempted by substitution of excised urethral segments with either homologous (rabbit) or heterologous (canine) urethral ECM grafts. Both groups demonstrated adequate tissue regeneration without evidence of scarring or stricture up to 8 months post-implantation. Tissue regeneration with homologous ECM resulted in more complete smooth muscle regeneration. In another investigation, rabbit bladder ECM or rabbit bladder ECM seeded with cultured epidermal cells was used as a urethral replacement graft. The cell-seeded group demonstrated more completed regeneration of mucosa and smooth muscle. The unseeded group had a higher stricture rate (Fu et al., 2007). Homologous genitourinary grafts have also been used to regenerate ureter and vagina. Ureteral regeneration was attempted in a rat model by creating a ureteral ECM graft and using this graft to replace 3–8 mm of excised ureter. The grafts were implanted over a polyethylene stent that was left in situ until the animal was killed. Similar to experiments using SIS for ureteral replacement, adequate regeneration was noted. Hydronephrosis was evident in cases where the stent migrated away from the graft. It is

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likely that all ureters would stricture to some degree following stent removal (Dahms et al., 1997). Vaginal replacement was evaluated in a series of rats that underwent hysterectomy and either primary closure of the vaginal stump, or grafting of the vaginal stump with vaginal ECM or bladder ECM. The grafted vaginas were similar in length to the pre-operative vaginas while the control vaginal stumps experienced significant contraction. Tissue regeneration was similar and no significant difference in vaginal length was noted for vaginas grafted with vaginal or bladder grafts. The grafts did undergo circumferential narrowing, which is consistent with all attempts to regenerate smalldiameter, hollow-lumen genitourinary organs (Wefer et al., 2002).

12.4.3 Acellular dermal matrix A majority of the natural ECMs utilized for tissue engineering in urology have been derived from visceral organs with a mucosal surface. Within other fields of medicine (general surgery, otolaryngology, plastic surgery), the most widely used ECM is derived from the dermis. Acellular dermal matrices are produced from a number of different species and by varied manufacturing processes. The basic components of dermal ECM are similar to the components of other ECMs. One important variation is that dermal ECM is primarily composed of type I collagen while intestinal and bladder ECM are composed of type I collagen and type III collagen (Chen et al., 2004). A limited number of investigations have utilized dermal ECM as scaffolds for tissue regeneration of genitourinary organs. Akbal et al. (2006) performed bladder augmentations in pigs using human acellular dermal biomatrix (AlloDerm®, LifeCell Corporation, Branchburg, NJ, USA). The authors cite preliminary data that suggested dermal ECM could result in regeneration of all of the components of the bladder wall. In order to test the results in pigs with outlet obstruction, the authors created a model of obstruction and used urodynamics to confirm poor compliance in all of the obstructed animals. The animals then underwent bladder augmentation with dermal ECM and relief of outlet obstruction. Three months following augmentation the animals again underwent urodynamics, which failed to demonstrate any improvement in compliance. Histological analysis at that time demonstrated poor tissue regeneration with significant fibrosis. The authors concluded that dermal ECM could not reliably produce tissue regeneration in a diseased animal model. A second investigation utilized porcine acellular collagen matrix (Pelvicol®, Bard Inc., Covington, GA, USA) to perform bladder augmentations in rabbits. Pelvicol®, unlike the other types of ECM used in urology, is chemically cross-linked making it a permanent implant. At the time of bladder harvest, up to 3 months post-implantation, gross and histological

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inspection demonstrated abscess and granuloma formation, microcalcification of the graft, and limited tissue regeneration (Ayyildiz et al., 2006). These results mimic findings encountered with other permanent implants and underscore the fundamental principle that natural biomaterials for tissue regeneration must be biodegradable in order to promote optimal tissue regeneration.

12.5

Future trends

This chapter focused on the fundamentals and experimental applications of natural biomaterials. The materials utilized in tissue engineering have been refined over time, yet the optimal biomaterial has not been identified. The limitations of tissue engineering, however, are not solely found in the biomaterials. The current understanding of the molecular pathways of regeneration is incomplete. Important research is beginning to identify significant components of the pathway. In the above examples it was observed that the urothelium is always the first portion of the bladder to regenerate. It is likely that the urothelium influences the remainder of the tissue components (i.e. smooth muscle) to regenerate in an orderly fashion. The molecular interaction between urothelium and bladder mesenchyme to produce bladder smooth muscle in embryonic mouse bladders is being intensely studied (Shiroyanagi et al., 2007). The molecular signals found in this developmental process will serve as a guide to similar pathways in tissue regeneration. Developmental biology also provides one true example of tissue regeneration within mammals. It has been observed that fetal wounding during early gestation results in healing without scar formation. Fetuses of various species have demonstrated the propensity to completely regenerate injured tissue, especially skin. Several factors, including elevated levels of hyaluronic acid within fetal tissue, have been suggested as a key difference between fetal tissues and tissues of post-gestational animals (Yannas et al., 2007). The model of early fetal healing will continue to provide evidence linking scarless wound healing and tissue regeneration. The future of natural biomaterials will likely include modifications to enhance the physical properties of the graft and to provide molecular promoters of regeneration. Investigators have recently explored the potential of enhancing bladder ECM grafts through the addition of hyaluronic acid to the grafts. Brown et al. (2006) provided evidence that smooth muscle cells cultured on bladder ECM enhanced with hyaluronic acid demonstrated improved smooth muscle cell contraction and proteolysis over cells cultured on unaltered bladder ECM. Separate researchers have investigated the physical alteration of SIS through the addition of polymer nanoparticles. Mondalek et al. (2008) have

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successfully altered the permeability of SIS through the addition of poly(lactic-co-glycolic) acid polymers. When considering the use of SIS for urinary tract regeneration, optimal permeability would allow tissue regeneration while limiting extravasation of urine and therefore reducing inflammation. Future applications of this technology may allow for covalent bonding of active molecular compounds to the nanoparticles. In this manner, both the physical and molecular properties of the graft could be optimized. The future of natural biomaterials will continue to broaden as the understanding of the molecular pathways of regeneration evolves.

12.6

Sources of further information and advice

All of the cited references provide insight into the development and application of natural biomaterials. Augst et al. (2006) offer an excellent overview of the biochemical properties of alginate and how these properties are utilized to create a fundamental biomaterial. Li (2000) discusses the relevant processing of naturally derived collagen and the biochemical properties of biomaterials produced from collagen. Badylak (2004) reviews the important components of ECM grafts and how these components contribute to tissue development and maturation. It is clear from the above discussion that biomaterials can be produced in a number of different ways. For example, the processing methods for bladder submucosa in one lab may differ from the methods utilized by a different group of researchers. Additionally, the production methods for the commercially available biomaterials are often not completely disclosed to the researchers utilizing the products and will evolve over time. While there is no one resource that will provide this information, researchers are urged to communicate with commercial manufacturers and other researchers to ensure that variations in experimental outcomes are not the result of altered biomaterials.

12.7

References

akbal c, lee sd, packer su, davis mm, rink rc, kaefer m (2006), ‘Bladder augmentation with acellular dermal biomatrix in a diseased animal model’, J Urol, 176(4 pt 2), 1706–1711. allman aj, mcpherson tb, badylak sf, merrill lc, kallakury b, sheehan c, raeder rh, metzger dw (2001), ‘Xenogeneic extracellular matrix grafts elicit a TH2restricted immune response’, Transplantation, 71(11), 1631–1640. atala a, kim w, paige kt, vacanti ca, retik ab (1994), ‘Endoscopic treatment of vesicoureteral reflux with a chondrocyte-alginate suspension’, J Urol, 152(2), 641–643. augst ad, kong hj, mooney dj (2006), ‘Alginate hydrogels as biomaterials’, Macromol Biosci, 6(8), 623–633.

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13 Nanotechnology and urological tissue engineering B. S. H A R R I S O N and C. L . WA R D, Wake Forest Institute for Regenerative Medicine, USA

Abstract: This chapter focuses on the possibilities of nanoscale tissue engineering strategies that are directed towards common urological problems, utilizing nanomaterials to replace or improve upon current methods. In general, nanomaterials are materials that possess at least one dimensionality between 1 and 100 nm and their small size creates powerful effects on cell biology. These materials also exhibit quantum phenomena, a potentially advantageous method for tissue engineering that has had little exploration. Three areas where nanomaterials are beginning to be used in tissue engineering will be investigated: biomaterials, cell tracking, and drug delivery. Key words: nanomaterials, cell tracking, drug delivery, quantum.

13.1

Introduction

Kidney, bladder, prostate, and other urological problems are a common occurrence in the world today. Such medical problems currently affect a large number of patients and consume a large amount of money in the health system. For example, over-active bladder syndrome affects 16–18% of adults (Milne, 2008), prostate cancer affects at least 18% of the male population (Wigle et al., 2008) and over 400 000 Americans have end-stage renal disease (Go et al., 2004). Current strategies to address these issues are not ideal because of increased cost, ineffective long-term replacement requiring additional surgeries, ineffective medicines, and other problems that hinder a full return to normal health. Tissue engineering may provide an answer to such issues, specifically focusing on the possibilities on a nanoscale. This chapter provides an introduction to nanomaterials and discusses the features that make them so relevant for use in tissue engineering. Following this, three areas where nanomaterials are beginning to be used in tissue engineering – biomaterials, cell tracking, and drug delivery – will be discussed. 281

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13.2

Rationale for nanomaterials in engineering tissue

Tissue engineering represents one of the newest innovations in modern-day science. It explores the possibility of growing biological substitutes to repair diseased or damaged tissue to restore normal function. The basic method of tissue engineering is that cells can be expanded in vitro, placed on a tissue scaffold made of a suitable biomaterial, and then implanted into the host. This paradigm has been demonstrated for several urological tissues. For example, a complete neobladder has been created in vitro and implanted into patients with spina bifida (Atala et al., 2006), polymer scaffolds have been created for urethra reconstruction (De Filippo et al., 2002), and embryonic stem cells have been shown to self-organize into glomeruli and tubulelike structures, secreting urine in a similar way to kidneys (Lanza et al., 2002). Other examples of tissue engineering principles being used in urological tissue have been reviewed elsewhere (Atala, 2003; Cross et al., 2003). While the proof of the concept of tissue engineering demonstrates its feasibility as a viable alternative in the treatment of severe urological disease, conventional engineering improvements in the materials used will allow greater advances in enhancing the success of engineered tissue. There are three areas where improvements in the materials can be made and then applied to tissue engineering: biomaterials, cell tracking, and drug delivery. First, improved biomaterials will remove the constraints currently faced in their use and better compliment the body’s natural regenerating process. Second, improvements in tracking cells will allow the development of better methods for evaluating the efficacy of cellular therapies. More sensitive and more robust imaging agents can more accurately measure where cells go after injection; more sensitive agents could be used to diagnose cancer or provide relevant physiological information in the microenvironment through less invasive means. Third, improvements in drug delivery methods will allow the reintegration of the engineered tissue within the body to occur more smoothly. For example, drug delivery devices could pinpoint the exact locations of diseased tissue and then exhibit controlled release of multiple types of drugs could be used to orchestrate the regeneration process. One approach to improving materials for urological tissue engineering is using nanomaterials. In general, nanomaterials are materials that possess at least one dimensionality between 1 and 100 nm. Their sizes and shapes can be easily controlled giving spheres, rods, squares, fibers, etc. Because of their small size one would expect these materials to have potentially powerful effects on cell biology. This should not be surprising since nanomaterials are comparable in size to proteins and viruses. Cells and the extracellular matrix (ECM) also possess a multitude of nanodimensionalities that interplay with one another. Cells, typically microns in diameter, are composed of numerous nanosized components all

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working together to create a highly organized, self-regulating machine. For example, the cell surface is composed of ion channels that regulate the passage of ions such as calcium and potassium in and out of the cell. Enzyme reactions, protein dynamics, and DNA all possess some aspect of nanodimensionality. These nanodimensional components control how cells produce the ECM including the ECM composition and architecture. The ECM that cells interact with also abounds with nanosize features that influence the behaviors of other cells and tissues. Nanosized features, such as fiber diameter and pores, along with the intrinsic properties of the matrix itself, control the mechanical strength, the adhesiveness of the cells to the matrix, cell proliferation, and the shape of the ECM. In addition to the fact that nanomaterials are of a size that is comparable to biological processes, quantum phenomena begin to appear when materials are nanosized. These quantum effects could be exploited to provide new approaches to regenerative therapies. Such quantum effects result in high optical absorptivities coupled with large photostabilities, or unusually magnetic properties within nanomaterials. Such nanomaterials are already being explored to enhance cellular imaging (Zhang et al., 2002; Medintz et al., 2005). Besides imaging, these quantum effects will allow novel methods of drug delivery using light, and electric or magnetic fields as drugdelivering triggers. These quantum effects are one of the most exciting reasons why nanomaterials should be explored for use in tissue engineering. Three areas where nanomaterials are beginning to be used for biomedical applications such as tissue engineering will now be examined. The three areas to be explored are the use of nanomaterials as biomaterials, for cell tracking, and for drug delivery.

13.3

Use of nanomaterials as biomaterials

A biomaterial in its most simplistic manifestation is the insoluble matter that provides the support structure for cells. They are particularly important to use if one is replacing a large tissue mass. However, biomaterials serve not only to define the three-dimensional space that the cells occupy, but also help control the tissue dynamics within the microenvironment. Thus, in order to control cell behavior, it is important to have control over the properties and use of biomaterials. There are several ideal characteristics of biomaterials. Ideally, the biomaterial should be biodegradable and bioresorbable. The biomaterial should provide the necessary mechanical strength to support the engineered tissue into the early stages of implantation and should not hinder development of new tissue growth (Babensee et al., 1998). It should minimize inflammation, and support its eventual replacement by the native ECM. Biomaterials have been traditionally divided into two major classes: natural and synthetic

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biomaterials. Each class of biomaterial has its distinct advantages and disadvantages. Naturally derived (e.g. collagen and alginate) and acellular (e.g. bladder submucosa and small intestinal submucosa) biomaterials have been extensively used in tissue engineering. These biomaterials are similar in structure to natural ECMs and may already possess the three-dimensional structure and functional proteins, thereby aiding in constructive tissue remodeling instead of scar tissue formation (Badylak, 2004). However, they are expensive to produce on a large scale and batch-to-batch variations can make it difficult to obtain reproducible materials. Synthetic materials such as poly-l-lactide acid, poly-caprolactone, and polyurethanes have been extensively used in tissue engineering. A major drawback is that many current synthetic materials lack the structural and chemical cues that many naturally derived materials possess. While synthetic polymers may be oversimplified in terms of structure compared with naturally derived materials, they are relatively easy to synthesize, to tune their physical properties, and to process into tissue scaffolds. In many cases synthetic polymers are advantageous when high mechanical strength or long-term stability is needed. However, a significant disadvantage is that changes in the chemical structure may affect several properties at one time. One possible way for materials to be more conducive to cell growth and proliferation is controlling the nanomorphology. This modification can be done by chemical or physical means. Chemical methods typically focus only on changes to the surface layer of the biomaterials without changing the bulk properties. This allows desirable properties to be retained, such as mechanical strength or biodegradation rate, while making the material more amendable for cells. Physical methods are principally used to control the nanomorphology and typically change the porosity or density of the materials. In order to illustrate how modifying the nanomorphology on the surface chemically affects cells, one can examine the use of scaffolds for bladder regeneration. Cells respond to the substructure of a surface and this drives their adhesion ability, attachment strength, and phenotypic expression. The correct nanostructure of a biomaterial that will be used for bladder regeneration may allow better cellular responses to the substrate, allowing the cells to adhere to and proliferate on the biomaterial. The scaffolds are typically constructed from a synthetic polymer and modified chemically to define the nanostructure characteristics. Studies have shown that polylactic-co-glycolic-acid (PLGA), poly-ether-urethane (PU), and polycaprolactone (PCL) can be treated with sodium hydroxide (NaOH) and nitric acid (HNO3) to increase the surface area and roughness as the surface feature dimensions are reduced to the nanoscale level (Thapa et al., 2003a). The modifications create cell-binding sites for adhesion, which are not normally present on the unmodified synthetic scaffolds. Smooth muscle cells

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have been shown to adhere successfully to the nanostructure and are able to exhibit long-term growth and proliferation (Thapa et al., 2003a). Similar scaffolds have been tested in vivo as patches for bladder reconstruction with some success (Pattison et al., 2007). Another example where controlling the nanoscopic features chemically has an impact on tissue engineering is in the area of stem cell differentiation for engineered tissue. Currently, concoctions of expensive growth factors are used to guide the differentiation of stem cells down certain lineages. With the ability to control the surface morphology and chemistry at the nanoscale, nanomaterials may eliminate the need to culture different cell types for reassembly into an engineered tissue as they can recruit the body’s own stem cells and differentiate them into the correct phenotype. Neural progenitor cells have been shown to differentiate by the presence of nanoscopic features. Networks of nanofibers were formed by self-assembly of peptide amphiphile molecules and cell suspensions, presenting the cells with a specific neurite-promoting laminin epitope. This induced differentiation of the cells into neurons (Silva et al., 2004). Another chemical modification can occur on the surface, using self-assembled monolayers to define the nanostructure (Liu et al., 2000). The potential of surface alterations at nanoscale dimensions have presented promising results for urological tissue engineering. Cells thrive and function is improved on these surfaces, as compared with a featureless polymer substrate; therefore, these results represent the beginning of more advanced biomaterials and enhancement of current medical techniques. Biomaterials can also be physically formed or molded into structures containing nanofeatures. Examples of some techniques used for creating nanostructured surfaces for tissue engineering are shown in Table 13.1. Of

Table 13.1 Examples of tissue scaffolds created using nanofabrication techniques Technique

Tissue scaffold prepared

Lithography Electrospinning

Nerve (Gabay et al., 2005) Heart (Zong et al., 2005) Nerve (Yang et al., 2005) Bone (Fujihara et al., 2005) Nerve (Ellis-Behnke et al., 2006) Bone (Du et al., 1999, Kikuchi et al., 2001, Kim et al., 2006, Liao et al., 2004) Bladder (Pattison et al., 2005, Thapa et al., 2003a, Thapa et al., 2003b) Bladder (Pattison et al., 2005, Thapa et al., 2003a, Thapa et al., 2003b)

Self-assembly Polymer demixing Solvent casting Salt leaching

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the many techniques available, there is a trade off between control of nanofeatures and cost or the level of expertise required to perform the techniques. For example, salt leaching and solvent casting methods are relatively simple; however, regular control over the nanostructure is difficult. In contrast, lithography, which gives a high degree of control, is expensive and requires a high level of expertise to perform. Electrospinning, however, is an attractive method for forming nanostructure biomaterials for tissue engineering. Electrospinning is a process where a fiber is produced from a solution using electrostatic forces. Briefly, a polymer solution is created, typically by dissolving a polymer in an organic solvent. The solution is placed in a syringe with a needle-tip and placed on a syringe pump that can expel the solution at a constant rate. A high voltage potential is applied to the polymer solution inducing free charges into the solution. A grounded collector is placed in front of the solution. The free charges in the solution respond to the applied electric field towards the electrode of opposite polarity, or the grounded collector. At the tip of the needle, the solution is forced out in a cone-like projection in the presence of the electrical field. When the applied potential reaches a critical value that overcomes the surface tension of the liquid, a stream of liquid is ejected from the tip. After this ejection, the stream experiences chaotic motion and bending instability as it collects on to the grounded surface. The travel distance through the atmosphere allows for solvent evaporation, resulting in a dry fiber that is collected on the surface. Depending on several factors (such as collector distance, polymer solution viscosity, applied potential, etc.), the fiber diameter can be controlled. The average diameter range is between 100 and 500 μm (Matthews et al., 2002). This process offers a suitable fabrication technique for the creation of nanostructured features. Biomaterials play an important role in regenerative medicine with their use in implants and tissue scaffolds. Through the ability to control the nanostructure of a biomaterial, better understanding and control of cell behaviors will result, creating better regenerative therapies. The impact of nanotechnology on biomaterial development, as it relates to regenerative medicine, will first be felt through better performing, passive tissue structure and will eventually progress to smart biomaterials that can be implanted and can then direct the regenerative process at the cellular level. Each type of biomaterial has its advantages and disadvantages and must be selected based on the needs of the engineered tissue (Ratner and Bryant, 2004). The advantages and disadvantages have to be carefully weighted and considered, especially since the material requirements for tissue engineering exceed those of traditional medical devices.

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13.4

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Use of nanomaterials for aiding cell tracking

A second area where improvements in tissue engineering are needed is tracking cells. Cellular therapies have been limited by a lack of adequate tracking methods to better understand where cells go in the body. For example, magnetic resonance imaging (MRI) is currently used to diagnose disease. It has high sensitivity when diagnosing the absence of cancer, but it has been shown to have very low sensitivity when positively detecting prostate cancer, which leads to a low accuracy in diagnosis (Lowe and Nagler, 2005). This illustrates a similar problem in tissue engineering where the ability to monitor cells or tissue scaffolds non-invasively once they are placed in vivo is limited. Labeling of cells with organic dyes or tracers for cell imaging has several drawbacks, particularly for in vivo imaging. Labeled cells cannot be tracked for very long because of poor photostability and low optical absorptivity of the organic dye. Dyes can also be chemically degraded, which shortens their lifetime. Background autofluorescence is difficult to distinguish from the short fluorescent lifetimes common to organic dyes. Finally, optical absorption of surrounding tissue makes it difficult to measure fluorescence intensity through the tissue. These are just a few reasons for the limitations of cell tracking, thus necessitating the need for newer materials. Nanomaterials are beginning to be created that are able to track cells and even assess the physiological state of the area in which they are placed. In vitro and in vivo visualization of nanoscale systems can be carried out using a variety of clinically relevant modalities such as fluorescence microscopy, single photon emission computed tomography (SPECT), positron emission tomography (PET), MRI, ultrasound, and radiotracing such as gamma scintigraphy. Nanoparticulate imaging probes include semiconductor quantum dots, magnetic and magnetofluorescent nanoparticles, gold nanoparticles and nanoshells among others. While there are currently few examples of nanomaterials being applied to the understanding of important processes in tissue regeneration, relevant uses of nanoparticles for regenerative medicine such as monitoring angiogenesis (Winter et al., 2003) and apoptosis (Jung et al., 2004) are appearing. Quantum dots are one type of nanomaterial that is receiving special attention. Quantum dots are inorganic nanocrystals that possess physical dimensions between 2 and 10 nm. The emission wavelength is controlled by the size of the nanocrystal and can be tuned throughout the visible spectrum to the near-infrared region (>670 nm). Early live cell experiments using fluorescent quantum dots sparked interest in using nanoparticles for immunocytochemical and immunohistochemical assays as well as for cell tracking (Akerman et al., 2002, Tokumasu and Dvorak, 2003, Sukhanova et al., 2004). A significant advantage for quantum dots is their increased

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photostability (typically 10–1000 times more stable) compared with organic dyes. This allows quantum dots and the cells or proteins attached to them to be tracked over longer periods of time. Tumor cells labeled with quantum dots have been intravenously injected into mice and successfully followed using fluorescence microscopy (Gao et al., 2004; Voura et al., 2004). As passive imaging agents, quantum dots can be used for imaging microvascularity in animals since poly(ethylene glycol) (PEG)-coated quantum dots injected into mice have shown good tissue perfusion and appear to be biocompatible (Ballou et al., 2004). Quantum dots represent just one novel class of nanomaterials whose ability to aid in the long-term imaging of cells would help develop better regenerative therapies. Other nanoparticles are showing promise for optical cell tracking and imaging. For instance, nanosized tubes of carbon, known as carbon nanotubes, possess optical transitions in the near-infrared region (>670 nm) that can be utilized for tracking cells. The infrared spectrum between 900 nm and 1300 nm is an important optical window for biomedical applications because of the lower optical absorption (greater penetration depth of light) and small autofluorescent background. Like quantum dots, carbon nanotubes possess good photostability and can be imaged over long periods of time using Raman scattering and fluorescence microscopy. However, unlike quantum dots, which are typically composed of heavy metals such as cadmium, carbon nanotubes are made of carbon, an abundant element in nature. Carbon nanotubes possess large aspect ratios with nanometer diameters and lengths ranging from submicron to millimeters. These tubes can contain a single wall of carbon or multiple walls (typically three to ten) of carbon, commonly called single-wall carbon nanotubes (SWNTs) or multiwall carbon nanotubes (MWNTs), respectively. SWNTs dispersed in a Pluorinc surfactant can be readily imaged through fluorescence microscopy after being ingested by mouse peritoneal macrophage-like cells. The small size of the SWNT makes it possible for 70 000 nanotubes to be ingested where they can remain stable for weeks inside 3T3 fibroblasts and murine myoblast stem cells (Cherukuri et al., 2004, Heller et al., 2005). Having such a high concentration of carbon nanotubes within a cell without distributing the cell behavior means such probes could be used for studying cell proliferation and stem-cell differentiation, even through repeated cells. While such nanomaterials have yet to reach clinical applications, it does show the potential for non-invasive optical imaging. Along with optical contrast agents, magnetic nanoparticles have also been used to track cells and report on cell behavior. Many nanoparticle contrasting agents are based on superparamagnetic iron oxide nanoparticles and some have already been approved as clinical MRI contrast agents. When placed into a magnetic field, magnetic nanoparticles create perturbations of the external field that significantly reduce the spin–spin relaxation

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time (T2) of the nearby environment generating MR contrast. Typically, these probes consist of a magnetic iron oxide core that is surrounded by a biocompatibilizing material such as dextran. The sizes of these particles can range from 1 nm to hundreds of nanometers in diameter. When used in conjunction with HIV-Tat (a cell-permeable peptide) and poly-arginine peptides, these particles are readily taken up by many cell types (Dodd et al., 2001; Zhao et al., 2002). For example, superparamagnetic iron oxide (SPIO)-labeled rat mesenchymal stem cells injected into rats could be imaged and tracked to the liver and kidneys (Bos et al., 2004). Apoptosis is commonly detected by using the binding of annexin V to externalized phosphatidylserine. This binding event is the basis of optical and radiolabeling methods for detecting apoptotic cells and can be bound to iron nanoparticles for sensing using MRI. It has been demonstrated that tumor-bearing mice injected with SPIO particles bearing apoptotic sensing proteins showed a sharp decrease in the T2* (the spin–spin relaxation time with the added inhomogeneous nature of the magnetic field) weight image corresponding to the location of the tumor (Zhao et al., 2001). This proved that nanomaterials can be used to create high-specificity MRI contrast agents for apoptotic cells. Such results are encouraging because they show that nanomaterials can be used not only for imaging the physical location of cells but also for providing information on the biological state of cells. While MRI has revolutionized our way of visualization in vivo, allowing cells to be tracked non-invasively, it is difficult to quantify the MRI signals and provide real quantification of cell numbers. The difficulty arises because MRI contrasting agents that are based on paramagnetic gadolinium and iron metals are not directly detected by the scanner but are indirectly detected by their influence on surrounding water molecules. However, the use of perfluoronated nanoparticles has recently been shown to be a new way to provide quantitative numbers to MRI since the fluorine nuclei (19F) can be directly detected (Morawski et al., 2004; Ahrens et al., 2005). Since endogenous fluorine is negligible in the body, 19FMRI is capable of directly detecting fluorine against a dark background similar to radiotracers and fluorescent dyes. While this has been demonstrated with dendritic cells, similar results should be obtainable using other cell types. As nanotechnology progresses new nanomaterials and techniques are being developed regarding cellular imaging that will better equip those practicing tissue engineering to reach their goals. Cellular therapies for regenerative medicine would benefit from nanotechnology since tracking of implanted cells would provide the means to better evaluate the viability of engineered tissues and help in understanding the biodistribution and migration pathways of transplanted cells. Nanotechnology would also allow better and more intelligent control of the bioactive factors that can influence cellular therapies. The potential impact of nanotechnology on

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regenerative medicine is great, creating the hope for individualized and targeted therapies.

13.5

Use of nanomaterials to improve drug delivery

Techniques for achieving adequate delivery of medicine and drugs to areas of the body have been explored for some time. An ideal treatment would be applying a drug to the required area so that it can be concentrated to the specific target and have no adverse effects on surrounding tissues. Current treatments are usually given intravenously or orally, which disperses the drug throughout the body, decreasing the concentration and applying it to unnecessary regions. Nanotechnology has the capability to produce carriers for delivery of therapeutics in order to aid specific treatments in tissue engineering. For example, biodegradable nanoparticles can deliver drugs, growth factors, and other bioactive agents to cells and tissue (Panyam and Labhasetwar, 2003). Nanodelivery vehicles possess three distinct advantages over conventional drug delivery methods. First, nanoparticles, owing to their small size, are able to bypass biological barriers such as cell membranes and the blood–brain barrier, allowing greater concentrations of therapeutics to be delivered. Second, nanocarriers can be functionalized with active targeting agents to allow selective delivery of bioactive agents. Third, drug delivery systems can incorporate nanotriggers for noninvasive delivery of therapeutic agents. These sensitive triggers can be activated using in vivo signals such as pH, ion concentration, and temperature or external sources such as near-infrared light, ultrasound, and magnetic fields. A summary of the features of the a nano-based drug carrier is given in Table 13.2. In several urological applications, nanoparticles have performed as suitable delivery vehicles for drugs and therapies. Prostate cancer is a common target for such drug delivery. Studies have shown that chemotherapy drugs, such as paclitaxel, can be loaded into nanoparticles to inhibit human prostate cancer cell lines (Sahoo et al., 2004). Therapeutic agents like oligonucleotides can also be formed into nanoparticles. The cells can uptake these agents in the presence of polypropylenimine dendrimers which suggests that the particles have the capability of delivering therapeutics to metastatic prostate cancer cells in the body (Santhakumaran et al., 2004). Similarly, specific molecule delivery can be addressed by targeting synthesized antibodies against prostate-specific membrane antigens with conjugated dendrimer nanoparticles (Thomas et al., 2004). Gene therapy is also an advancing field with nanotechnology, manipulating DNA on a nanoscale. Novel folatelinked nanoparticles have been shown to increase the efficiency of in vitro DNA transfection (Hattori and Maitani, 2004). In addition, a human transferrin-targeted cationic liposome–DNA complex, known as

Table 13.2 Components of a nano drug delivery system (NDDS) Binder

Biocompatibilization

Imaging contrast

Sensor

Targeting

Therapeutics

The central component of the NDDS. All components of the NDDS are connected or attached to the binder. It may or may not play additional roles in the NDDS. The binder may also be the image contrasting agent, for example iron nanoparticles and quantum dots. Polymers such as polyglycolic acid may serve as the binder of the therapeutic agent and also the biocompatabilizing agent. This component makes the nanocarrier compatible with the biological environment. It does this by minimizing aggregation of the NDDS and increases the circulating lifetime of the carrier by avoiding the defense mechanisms of the biological systems such as the reticuloendothelial system. This component provides the means for imaging modalities to observe the nanocarrier. These contrasting agents may be observed using optical, magnetic, ultrasound, and scintillation methods. The sensor or trigger is used to alter the behavior of the nanocarrier once it has been deployed. For example, near-infrared light or electromagnetic radiation may be used to accelerate the release of a therapeutic or cause rapid localized heating as part of a therapy. Chemical sensors such as polymers that are pH or ion sensitive may also provide feedback to the nanocarrier in delivery of its payload. This component provides the means of driving the nanocarrier to its desired location. There are two types of targeting: passive and active. Passive targeting incorporates only non-specific targeting agents that may be useful for determining microenvironment permeability or areas of increased angiogenesis. Active targeting uses ligands or antibodies that bind to specific receptors at the target site. Active targeting aids in obtaining higher concentrations of therapeutics and contrasting agents at the desired site. In addition, multiple targeting agents can be bound to the nanocarrier allowing lower-binding-affinity molecules to be used to increase binding probabilities. Bioactive agents such as drugs or DNA are typical payloads of the nanocarrier. Drugs that are incapable of penetrating cellular membranes, or hydrophobic drugs that cannot be administrated systemically by themselves, can be contained within the nanocarrier awaiting release in a controlled manner. Other novel properties of nanoparticles have also shown promise as hyperthermic agents.

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transferrin-lipoplex, has demonstrated increased stability with enhanced in vivo gene-transfer efficiency and effectiveness for p53 gene therapy when it is used in sequence with common radiotherapy (Xu et al., 2002). Specialized drug delivery can offer tissue engineering a means of targeting particular areas in the body without damaging or even affecting surrounding tissues. Oral medications and injections tend to disperse in areas where they are not required, while nanosized drug applications can direct a concentrated amount of drug to the exact region of need: a benefit to patients suffering from side effects.

13.6

Conclusions

As nanotechnology continues to grow, it will provide new and powerful tools that will revolutionize regenerative medicine. The most significant impact that nanotechnology will have on regenerative medicine is that it will help in providing a detailed understanding and control of biology. Already, nanotechnology, albeit a young technology, has demonstrated significant advances over traditional imaging, sensing, and structural technologies. Many of these advantages stem from the capability of nanomaterials to be multi-functional. These advances help in tackling one of the most significant challenges we face in designing new biomedical technologies – targeting biological functions while at the same time avoiding non-specific effects. While there have been challenges for some time, nanotechnology provides us with the means to successfully negotiate these challenges and create new innovations in regenerative medicine.

13.7

Future trends

Nanomaterials are being embraced as the basis of the next generation of biomaterials for tissue engineering. With the exciting developments of new materials leading to more functional and higher performing biomaterials, three trends appear to be emerging. The first trend is the creation of smart nanomaterials, the second trend is related to the safety of nanomaterials, and the third trend is related to manipulating nanomaterials. Smart nanomaterials are part of a new class of multifunctional biomaterials. Unlike passive structural biomaterials, smart biomaterials are designed to interact actively with their environment either by responding to changes in their surroundings or by stimulating or suppressing specific cellular behavior. They can change their shape, porosity, or hydrophilicity based on changes in temperature (Gan and Lyon, 2001), pH (Bulmus et al., 2003), or external stimuli such as electric (Lahann et al., 2003) or magnetic fields (Jordan et al., 1999). Such control of the biomaterial behavior through nanotechnology could create a major shift in the way one uses biomaterials.

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Ideally, one would want to implant a biomaterial directly into the patient and the biomaterial would then selectively recruit the correct cell types to the correct location in the tissue. This method would be especially important for organs with very elaborate structures. A smart biomaterial would allow the correct cells and supporting vascular to grow on to the scaffold in the correct orientation without permitting inflammatory cells and fibroblasts, which typically wall off any implants, to become established on the biomaterial. Materials safety is of paramount concern whenever a new material is being used. Because nanomaterials exist on the same scale as viruses and other potent bioactive agents, it is important to understand how the materials interact with cells. Studying the cytotoxicity of carbon nanomaterials such as carbon nanotubes and buckyballs has yielded interesting and at times conflicting results. Some studies have indicated that carbon nanotubes are cytotoxic (Jia et al., 2005, Ding et al., 2005, Maynard et al., 2004, Shvedova et al., 2003). However, carbon nanotubes have also been reported to improve neural signal transfer and support dendrite elongation and cell adhesion in vitro (Lovat et al., 2005), as well as supporting other cell types such as smooth muscle (MacDonald et al., 2005), fibroblasts (Correa-Duarte et al., 2004), and osteoblasts (Supronowicz et al., 2002). This suggests that a combination of several factors are responsible for the observed effects (Nimmagadda et al., 2006). One observation is that the existing testing methods for cytotoxicity may not be sufficient to evaluate the affect of nanomaterials and thus new testing procedures need to be developed. Because of their small size, manipulating and assembling nanomaterials is challenging. Ongoing advances in surgical and robotic technology are emerging to provide automated mechanical devices with dimensions comparable to bacteria. The hope is that microscopic robots will be able to deliver drugs or target specific areas (Murphy et al., 2006). Current research has stepped into this horizon with the world’s smallest, controllable robot, measuring 250 × 100 μm (Murphy et al., 2006). The device has been shown to operate at distances greater than 35 cm and is controlled through teleoperation to navigate intricate paths, such as those in the organ systems. The small size would be able to address urological issues, such as cancer therapies, at a nanoscale level by implanting the device and controlling the required procedures from outside of the body. Another robotic application in nanotechnology is the Enseal® device (Tsiouris et al., 2006). The product is a surgical sealing device that deposits tissue-controlled energy on the nanometric scale for sealing and strengthening of vessels. After the application, vessel walls can withstand multiple times their normal pressure, which is beneficial to the dorsal venous complex in laparoscopic prostectomy and robotic-assisted laparoscopic prostectomy.

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13.8

Source of further information and advice

Additional information about nanomaterials and their use as biomaterials can be found in the Handbook of Nanostructured Biomaterials and Their Applications in Nanobiotechnology, edited by Hari Singh Nalwa and published by American Scientific Publishers.

13.9

References

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Molecular Imaging of angiogenesis in nascent vx-2 rabbit tumors using a novel alpha (v)beta (3)-targeted nanoparticle and 1.5 tesla magnetic resonance imaging. Cancer Research, 63, 5838–5843. xu, l., frederik, p., pirollo, k. f., tang, w. h., rait, a., xiang, l. m., huang, w., cruz, i., yin, y. and chang, e. h. (2002) Self-assembly of a virus-mimicking nanostructure system for efficient tumor-targeted gene delivery. Human Gene Therapy, 13, 469–481. yang, f., murugan, r., wang, s. and ramakrishna, s. (2005) Electrospinning of nano/ micro scale poly (L-lactic acid) aligned fibers and their potential in neural tissue engineering. Biomaterials, 26(15), 2603–2610. zhang, y., kohler, n. and zhang, m. q. (2002) Surface modification of superparamagnetic magnetite nanoparticles and their intracellular uptake. Biomaterials, 23, 1553–1561. zhao, m., beauregard, d. a., loizou, l., davletov, b. and brindle, k. m. (2001) Noninvasive detection of apoptosis using magnetic resonance imaging and a targeted contrast agent. Nature Medicine, 7, 1241–1244. zhao, m., kircher, m. f., josephson, l. and weissleder, r. (2002) Differential conjugation of tat peptide to superparamagnetic nanoparticles and its effect on cellular uptake. Bioconjugate Chemistry, 13, 840–844. zong, x. h., bien, h., chung, c. y., yin, l., fang, d., hsiao, b. s., chu, b. and entcheva, e. (2005). Electrospun fine-textured scaffolds for heart tissue constructs. Biomaterials, 26(26), 5330–5338.

14 Assessing the performance of tissue-engineered urological implants G. J. C H R I S T, D. B U R M E I S T E R, S. V I S H WA J I T, Y. JA R A JA P U and K.- E. A N D E R S S O N, Wake Forest Institute for Regenerative Medicine, USA

Abstract: Regenerative medicine and tissue engineering technologies are rapidly advancing with potentially broad clinical applications, including those in the field of urology. Although all structures of the lower urinary tract (ureters, bladder, sphincter, urethra) would undoubtedly benefit from utilization of tissue engineering technologies, this report will focus on applications in the bladder. In this regard, the results of both preclinical and clinical investigations support the use of engineered bladder constructs/implants for the treatment of end-stage bladder disease. Despite great progress, further optimization and utilization of this groundbreaking technology will require a more complete characterization and understanding of bladder regeneration both in vitro and in vivo. The complexity of normal tissue structure and physiology, in turn, requires a rigorous examination of the engineered ‘biomimetics’ to ensure that they do indeed provide the required structure, physiology and function. In this chapter we outline an algorithm for evaluation of bladder function that will be useful for such a characterization of the physiological attributes of engineered and regenerating bladders. Briefly, we propose a multidisciplinary ‘vertical approach’ that assesses the characteristics of bladder function/phenotype at the genetic, cellular, molecular, tissue and whole animal level. Appropriate methods and instruments are already available to this end, and in this chapter we describe both the methods and rationale for comparison of engineered/ regenerating bladders with native bladder. Key words: bladder, physiology, function, tissue engineering, bladder regeneration, regenerative medicine.

14.1

Introduction

The aim of tissue engineering and regenerative medicine is the repair and replacement of damaged cells, organs and tissues. This important field of medical biology is rapidly evolving, and the degree of scientific sophistication and the concomitant technology array that can be applied to the field are also dramatically escalating. An ultimate goal must be to engineer 299

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tissues that approximate the native tissues they will replace as closely as possible. However, the complexity of normal tissue structure and physiology requires a rigorous examination of these engineered ‘biomimetics’ to ensure that they do indeed provide the required structure, physiology and function. With respect to this chapter, we are focusing our attention on the urological applications of tissue engineering/regenerative medicine. More specifically, the discussion herein is most pertinent to the treatment of advanced or end-stage bladder disease. In this regard, end-stage bladder disease affects thousands of patients with debilitating consequences, and can result from a variety of prevalent conditions (e.g. congenital malformations, nervous system conditions, cancer, infection and inflammation trauma) with a final common pathway ultimately leading to bladder failure or the need for bladder replacement. Bladder failure, in turn, can lead to compromised renal function and/or renal failure. Current management strategies for lower urinary tract reconstruction, of all types, involve use of gastrointestinal segments as tissues for lower urinary tract replacement or repair. However, gastrointestinal tissues are designed to absorb specific solutes, whereas bladder tissue is designed for the excretion of solutes. When gastrointestinal tissue is in contact with the urinary tract, multiple complications may ensue – such as infection, metabolic disturbances, urolithiasis, perforation, increased mucous production, and malignancy (Atala et al., 1993, Buscarini et al., 2007, Cabello Benavente et al., 2006, Farnham and Cookson, 2004, Gupta et al., 2007, Jensen et al., 2006, Kaefer et al., 1998, Koie et al., 2006, McDougal, 1992, Nabi et al., 2005, Parekh and Donat, 2007, Salle et al., 1990, Simon et al., 2006, Studer et al., 2006, Sullivan et al., 1998). Because of the problems encountered with the use of gastrointestinal segments, numerous investigators over the last 100 years have attempted to use alternative methods, materials and tissues for lower urinary tract replacement or repair (Atala, 1998, Atala, 1999, Bellinger, 1993, Cartwright and Snow, 1989, Dewan and Stefanek, 1994, Gonzalez et al., 1995, Salle et al., 1990). The limited success of these methods and our previous experiences with bladder replacement in dogs (Oberpenning et al., 1999) and humans (Atala et al., 2006) have encouraged us, and others, to continue to develop technologies using autologous engineered structures to replace the bladder and other components of the lower urinary tract. In fact, various strategies for urological tissue engineering/regenerative medicine ranging from complete organ regeneration in vivo to total organ engineering in vitro have been forwarded. Many of these approaches have been described in reviews over the past 5 years (Atala, 2003, Atala, 2005, Atala, 2006a, Atala, 2006b, Beiko et al., 2004, Frimberger et al., 2006, Korossis et al., 2006, Sievert et al., 2007, Staack et al., 2005). However,

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regardless of the precise strategy that is employed, the goal is always the same, i.e. to restore normal tissue/organ function. Thus, one must be able to evaluate the success of the strategy being employed in an objective fashion, in order to ensure that the engineered or regenerating tissue possesses the requisite phenotypic characteristics. Although urological applications of tissue engineering/regenerative medicine are being evaluated for their clinical utility throughout the lower urinary tract (ureter, bladder, sphincter, urethra, etc.), the largest database (both preclinical and clinical) is currently available for the bladder. As such, we will concentrate our discussion in this chapter on this tissue/organ. A brief description of bladder function and the unmet medical need is provided below.

14.2

The bladder

The primary function of the lower urinary tract is the unidirectional transport, storage and elimination of the urine produced by the kidneys. The lower urinary tract consists of several distinct parts: the bladder body (containing smooth muscle – the detrusor), the bladder base (neck), the trigone, the urethra and a specific uretero-vesical junction. The entire lower urinary tract is lined by a specialized urothelium, which covers the underlying lamina propria and smooth muscle cells. The bladder is thus the center point of the lower urinary tract, and a healthy, pain-free and continent bladder is important for the maintenance of a good quality of life. The normal function of the urinary bladder, in turn, is to fill, store and empty. That is, the bladder fills with urine produced by the kidneys, and is able to store increasing volumes at low physiological pressures. At maximum volume (i.e. bladder capacity), the detrusor smooth muscle of the bladder wall contracts with sufficient force to overcome the outflow resistance (i.e. urethra) and empty the contents of the bladder completely. This simple act reflects the complex orchestration of neuronal, myogenic and urothelial control mechanisms, and perturbations in any of these can result in bladder dysfunction. While viable treatment options are available for a variety of bladder diseases/dysfunctions, as described above, everyone recognizes the need for the development of more effective treatment options for advanced (i.e. end-stage) bladder disease and dysfunction. Interestingly, as outlined in Table 14.1, a review of the literature reveals that evaluation of the success/failure of bladder tissue engineering, to date, has focused largely on descriptive studies, mainly at the histological level. A few reports have even included in vivo bladder function studies (i.e. cystometry), while fewer still have performed physiological/ pharmacological evaluation of detrusor strips as well (Baumert et al., 2007, Campodonico et al., 2004, Cheng et al., 2007, Chung et al., 2005, Drewa et al., 2006, Farhat et al., 2006, Gilbert et al., 2005, Han et al., 2007,

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Table 14.1 Review of the bladder tissue engineering literature Reference

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Functional analysis

Han et al. (2007) Yu et al. (2007)

BAMG seeded with UCs implanted in nude rats Collagen-coated polycaprolactone scaffolds in rats

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Lai et al. (2006)

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BAMG seeded with UCs and SMCs implanted into pigs PGA scaffolds seeded with mouse fibroblasts implanted in rats Rabbit BAMG implanted into rabbits

Histology revealed presence of a multilayered epithelium Cystometry showed normal bladder capacity and micturition pressure Radiography showed a normal pear-shaped bladder Organ bath studies showed response to cholinergic stimulation Histology revealed a multilayered epithelium and differentiated SMCs Cystometry showed improved bladder capacity with seeded scaffolds Organ bath studies produced decreased response to EFS compared with native tissue Histology revealed improved SMC bundle organization with seeded scaffolds Cystometry demonstrated that storage function was retained Radiography showed good distension in BAMG-augmented rats Cystometry showed a decrease in bladder capacity with unseeded grafts IHC revealed increased SMC bundles at graft edges Histology revealed that fibrin glue helps to enhance cellular organization

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Histology revealed that fibroblast seeding enhances urothelial cell growth MRI showed enhanced angiongenesis with VEGF present on scaffold Histology revealed lower fibrosis with VEGF Cystometry 8 weeks after augmentation showed similar function to preoperative levels Histology showed BMDSCs and bladder SMCs enhance bundle formation Cystometry showed improved bladder capacity compared with cystectomy alone RT-PCR showed decreased collagen expression, and increased expression of MHC, CK8 and CK19 in BMDSCseeded group Mechanical features (strength) studied by increased strain to produce failure

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Table 14.1 Continued Reference

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Nuininga et al. (2004)

Yokoyama et al. (2001b) Schoeller et al. (2001) Lu et al. (2005) Zhang et al. (2004)

Kanematsu et al. (2003)

Lai et al. (2002)

Vascularized prefabricated flaps seeded with UCs in rat model SIS seeded with MDCs studied in vitro SIS seeded with human UCs and SMCs implanted into nude mice BAMG supplemented with fibroblast growth factor implanted into rats

Polymer scaffolds seeded with normal and diseased human SMCs implanted into nude mice

SEM revealed UCs seeded as a monolayer on surface of scaffold, but no penetration of SMCs Histology revealed some evidence of bladder regeneration, and the section of intestine used as scaffold may influence calcification Histology revealed that the urothelium develops to near native tissue by day 20 in a subcutaneous implant Histology revealed inflammation in the regenerated bladders with 21 out of 56 animals developing stones

Histology revealed that the treatment of regenerating bladders with MDCs resulted in less immune reaction (IHC vs CD4 and CD8) Histology revealed a contiuous urothelial layer, with UC-seeded capsules improving survival Mechanical studies showed areal strain is increased with MSC seeding; MDC growth supported by SIS Histology revealed SMC α-actin and CK A1/A3-positive cells

Organ bath studies showed regenerated bladder contracted in response to depolarization (KCl), cholinergic stimulation and EFS which was blocked by tetrodotoxin Organ bath studies produced typical and similar contractions in response to implants from both normal and diseased groups Histology revealed no difference in SMC markers from diseased cellseeded implants

BAMG, bladder acellular matrix graft; UC, urothelial cell; SIS, small intestinal submucosa; SMC, smooth muscle cell; IHC, immunohistochemistry; BMDSC, bone marrow-derived stem cell; ECM, extracellular matrix; RT-PCR, real-time polymerase chain reaction; MHC, myosin heavy chain; CK, cytokeratin; EFS, electrical field stimulation; MDC, muscle-derived cell; SEM, scanning electron microscopy; VEGF, vascular endothelial growth factor; PGA, polyglycolic acid MRI, magnetic resonance imaging.

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Kanematsu et al., 2003, Kimuli et al., 2004, Kropp et al., 2004, Lai et al., 2002, Lai et al., 2006, Lu et al., 2005, Nuininga et al., 2004, Obara et al., 2006, Oberpenning et al., 1999, Schoeller et al., 2001, Wang et al., 2005, Wunsch et al., 2005, Yang et al., 2005, Yokoyama et al., 2001a, Yokoyama et al., 2001b, Yu et al., 2007, Zhang et al., 2004, Zhang et al., 2006). Despite measurable progress, the extant work is not sufficiently comprehensive to elucidate the events/mechanisms responsible for the initiation, development and progression of bladder regeneration. In this regard, an accurate and rigorous assessment of the performance characteristics and durability of engineered and regenerating tissue will be a key factor in the iterative process that leads ultimately to routine application of these technologies for a broad range of clinical needs.

14.3

Evaluation of engineered or regenerating tissues in vitro

In this scenario, the over-riding goal of this chapter is to review critical technologies and assays required for an accurate phenotypic characterization of engineered and regenerating tissues/organs. A suite of proposed techniques is summarized in Table 14.2. A brief description of how in vitro methods can be applied to the study and improvement of the bladder tissue engineering process is described below.

14.3.1 Evaluation of contractile and relaxation responses in vitro The detrusor smooth muscle cell is a primary parenchymal cell type in the bladder wall. Given the central importance of contraction and relaxation to bladder function, a detailed understanding of the mechanisms and regulation of detrusor contractility is a critical metric for evaluating progress in bladder tissue regeneration and engineering. Basic knowledge of receptors, signal transduction and ion channel physiology/pharmacology of the detrusor smooth muscle has been well established in the literature. Therefore, the characteristics of native bladder function are quite well established (Andersson and Arner, 2004, Andersson and Wein, 2004, Brading, 2006, Christ and Hodges, 2006, Drake, 2007, Fry et al., 2004a, Fry et al., 2004b, Hegde, 2006, Heppner et al., 2000, Heppner et al., 2001, Herrera et al., 2003, Herrera et al., 2005, Michel and Vrydag, 2006, Ruggieri, 2006, Thorneloe and Nelson, 2003, Thorneloe et al., 2008, Werner et al., 2007). The proposed approach proceeds logically from this vantage point. The intuitive starting point in the characterization process is physiological and pharmacological studies using traditional organ bath methods

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Table 14.2 Metrics for evaluation of the functional performance of engineered and regenerating tissues Assay/technique

Functional significance

In vivo imaging (microcomputed tomography, MRI, PET, etc.)

Non-invasive assessment of tissue/organ growth, shape, dimensions; establish metrics for success/failure of engineered or regenerating tissues Invasive and non-invasive measures of the ability of the engineered tissue/organ to execute required functions in vivo In vitro evaluation of the characteristics of the isolated organ or tissue; these studies capture the extent of recovery of important functional characteristics and phenotypic characteristics Evaluation of the mechanical characteristics of tissue such as viscoelastic or inelastic properties Evaluation of the tissue/organ architecture, presence and distribution of cells, cell-specific markers, matrix, nerves, blood vessels, etc.

In vivo monitoring (cystometry, Doppler, EMG, etc.) Physiology of isolated tissues (organ bath, patch clamp, calcium imaging) Biomechanical testing (stress, strain, stretch, creep, etc.) Histological analysis (immunohistocemistry, immunocytochemistry, histomorphometry, etc.) Molecular analysis (Northern, Western and Southern blots, proteomics, genomics, etc.) Summary analysis

Establishing molecular ‘finger print’ of engineered tissues/organs; do the molecular characteristics of the tissue/organ match the native phenotype/configuration? Go/No Go*

MRI: magnetic resonance imaging; PET: positron emission tomography; EMG: electromyography. * Results of experiments with these methods will determine whether or not the technology is worth pursuing.

(Fig. 14.1(a)). This represents a widely used method that can be effectively utilized for the detailed functional evaluation of isolated tissue strips from native and engineered organs. The organ bath simulates critical aspects of the in vivo environment to ensure optimum physiological functioning of tissues being studied in vitro. This technique involves suspending excised detrusor tissue strips between two points (one fixed and one connected to a force transducer) under an optimum tension (i.e. basal tension or pre-tension to ensure that the muscle strips are in the appropriate configuration for measuring contraction; see below), and then following an appropriate equilibration period, exposing the isolated tissue to different pharmacological agents; in this instance, generally speaking, muscarinic and purinergic agonists and antagonists, as well as electrical field stimulation (EFS; selective stimulation of the intramural nerves to release endogenous

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14.1 Classical pharmacological studies conducted in an organ bath apparatus. (a) A representative example of acquired data from an experiment in rat bladder strips. Representative contractile responses observed in rat bladder strips in response to (b) carbachol and (c) electric field stimulation (EFS).

neurotransmitter(s)). Changes in the tension developed in response to pharmacological treatments are recorded isometrically. As noted above, application of an appropriate pre-tension determines the ‘performance’ of a tissue strip and is generally dependent on the amount of cellularity/ muscularity/matrix of the tissue, in this case the bladder. For bladder, pretension is ≈1–1.5 g, but for tissues derived from engineered neo-bladders one must first determine the length–tension relationship to ensure proper muscle performance and, therefore, correct interpretation of the data. For these purposes, the length–tension relationship is determined by recording the magnitude of the contractile response developed following addition of a receptor (i.e. carbachol) or non-receptor-mediated (KCl at depolarizing concentrations) agonist. The response to the same pharmacological stimulus is then recorded over a range of pre-tensions to determine the maximal degree of stretch (pre-tension). Once this is determined, the contribution of physiologically relevant receptor-dependent cellular signal transduction pathways to the contraction of isolated detrusor strips can be examined with confidence. Figure 14.1(b) illustrates a representative example of tension recording in a rat bladder strip contracting to increased concentrations of carbachol, a non-hydrolyzable muscarinic receptor agonist.

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As well as the addition of exogenous pharmacological agents, another important aspect of the functional evaluation of native and engineered neo-bladder is determining the responses evoked by activation of intramural nerves of the bladder wall. This is accomplished by applying an electric field (EFS) of an appropriate strength and recording contractile responses evoked in tissue strips. When properly applied, EFS causes release of endogenous neurotransmitters from depolarized intramural nerves to elicit contraction or relaxation of isolated tissue strips. For example, contraction to EFS in bladder strips is a well-recognized and physiologically relevant response. The strength of the EFS applied to the tissue is based on studies in native tissue strips. Thus, this method can be readily applied to engineered or regenerating bladder tissue strips, to evaluate the extent of neuronal re-innervation. When correlated with simultaneous immunostaining with neural-specific markers this provides a powerful method to evaluate the extent and functional significance of re-innervation of engineered tissues, which will ultimately be a critical factor in the widespread application of this technology. Figure 14.1(c) shows a representative example of tension recording in a rat bladder strip contracting to EFS applied at increasing frequencies. Coupling EFS with the addition of appropriate pharmacological agents can help establish the status and characteristics of bladder tissue regeneration, as well as the contribution of various neurotransmitters (e.g. muscarinic and purinergic) to engineered bladder function.

14.3.2 Real-time fluorescence imaging of intracellular messenger molecules Rapid and sensitive methods are available to evaluate the function of a living tissue at the cellular level. Imaging or tracking the activation/ production of an intracellular messenger molecule in response to physiological stimuli provides direct evidence for a functioning tissue. Depending on the tissue of interest, one can choose the intracellular molecule to be evaluated. With respect to the bladder, for example, the intracellular concentration of calcium ions is a very sensitive parameter to evaluate the effect of receptor or non-receptor activation of detrusor smooth muscle. Quantitative determination of nitric oxide (NO) production is another useful measure to evaluate the contribution of NO to regulation of detrusor contractility. In any instance, target-specific fluorescent molecules are available to determine qualitatively, as well as quantitatively, the production of different intracellular molecules. Fura-2 is a cell-permeant fluorescent dye that produces Ca2+-dependent fluorescence (Tsien et al., 1985) and is widely used for quantitative determination of intracellular [Ca2+] in response to agonist stimulation. This agent has been used to assess the role of calcium

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Basal

Carbachol 1 μm

10 μm

14.2 Increase in the steady-state free intracellular calcium ion concentration as determined using Fura-2 fluorescence ratios in smooth muscle cells in rat bladder strips in response to carbachol.

mobilization/flux in cells in a variety of tissues. Figure 14.2 shows a representative example of quantitative changes in fura-2 fluorescence in smooth muscle cells in rat bladder strips activated by the muscarinic agonist carbachol. By monitoring both dynamic and steady-state changes in [Ca2+] in cells and engineered neo-bladder tissues, signal transduction can be evaluated in a relatively high throughput fashion. DAF-2 is another cell-permeant dye that produces fluorescence only in the presence of NO and is widely used for quantitative determination of NO production (Kojima et al., 1998) in different cell types.

14.3.3 Evaluation of trans-membrane ionic currents Another critical approach to the functional evaluation of the phenotype of smooth muscle is the study of trans-membrane ionic currents in individual cells and changes in those currents in response to cellular activation/ stimulation. In this regard, trans-membrane ion currents are critical modulators of cellular homeostasis and excitability, and all cell types have characteristic ‘electrophysiological signatures’. Specifically, the contributions of ion channels to bladder smooth muscle function are well established (Brading, 2006, Chow et al., 2003, Fry et al., 1998, Fry et al., 2002, Heppner et al., 2003, Herrera et al., 2000, Herrera et al., 2001, Karicheti and Christ, 2001, Sui et al., 2003, Sui et al., 2007, Thorneloe and Nelson, 2003, Thorneloe et al., 2008, Venkateswarlu et al., 2002, Wang and Christ, 2001, Wang et al., 2006). Thus, analysis of ionic movement across the cell membrane is another important aspect in the functional evaluation of engineered tissues. Electrical activity in detrusor myocytes can be evaluated using either the whole cell configuration (single cells or dual whole cell patch to study intercellular communication through gap junctions) or membrane patch configuration (to study the biophysical properties of individual ion channels or small numbers of membrane ion channels). In either scenario, one can evaluate the characteristics of ionic currents and identify different ion channels involved in this ionic movement in response to cellular activation in the presence and absence of

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Current (pA)

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14.3 A representative example of outward potassium currents recorded in freshly dissociated bladder smooth muscle cells.

known pharmacological blockers of these ion channels. Figure 14.3 displays a representative example of outward potassium currents recorded in bladder smooth muscle cells derived from rats. Again, coupling this electrophysiological data with molecular identification of the ion channels (i.e. Northern, Westerns, etc.) and their distribution in the bladder wall is another powerful experimental tool for evaluating the physiology and function of regenerating bladder tissue (see section below for more details).

14.3.4 Biochemical evaluation Several biochemical techniques can also be applied to provide information that sheds important mechanistic insight on observations obtained with all of the aforementioned techniques. Engineered organs are expected to express receptors, ion channels and enzymes that are responsible for carrying out the key physiological function of the organ of interest. The bladder is no different, and expression of receptors and ion channels should be evaluated by determining the mRNA encoding the protein using the polymerase chain reaction (PCR) technique or by determining the receptor or ion channel protein in a representative sample of the engineered organ by Western blotting. Radioligand binding studies can also be a valuable adjunct and are highly recommended studies for quantification of functional receptor/ion channel binding sites and their distribution in a given tissue. This latter technique involves using receptor/ion channel-selective agonists and

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antagonists and provides information about receptor/ion channel characteristics and binding properties. These techniques are tedious but are being widely used and will provide valuable information about the molecular characteristics of the functional phenotype of different cell types in the engineered tissues. Studying the enzyme activity is yet another way of providing functional evidence. Extraction of an active enzyme from engineered organ samples is also tedious but involves routinely used laboratory techniques and provides valuable supporting information. There is no doubt that genomic and proteomic technologies will eventually be routinely applied to the study of engineered neo-organs as well. By coupling this quantitative information with the powerful mapping and pathway analysis tools that are quickly becoming available, one can clearly define the molecular events associated with normal (and by definition abnormal) bladder regeneration.

14.4

Bladder tissue engineering and regeneration

To emphasize the utility and importance of the proposed approach we present examples from our recent work on the bladder. Again, the choice of the bladder is based on the fact that we are able to apply the ‘vertical approach’ to our systematic analysis of engineered or regenerating tissue function (see Fig. 14.4), and moreover, that smooth muscle is a primary parenchymal cell type in the bladder making this tissue readily accessible to characterization with a host of well-established in vitro and in vivo methods/technologies. More specifically, by using the bladder as our model system we can conduct studies in vivo and then harvest the engineered bladder (as well as the native tissue) and conduct studies in vitro, at the cellular, molecular and genetic levels. This research approach provides a powerful experimental paradigm for evaluating the physiology of engineered and regenerating tissues/organs. To this end, we are currently pursuing two avenues of research designed to provide improved treatments for end organ bladder failure, and both provide excellent examples of how to assess the performance of engineered and regenerating tissues/organs. The first is bladder regeneration in vivo, and the second is bladder tissue engineering in vitro, prior to implantation and in vivo bladder regeneration. Despite the relatively straightforward function of the bladder, there is significant intrinsic complexity to bladder structure and function. Nonetheless, our focus herein is to describe the methods and technologies that can be used to assess bladder function and detrusor muscle contractility. Relevant animal models provide the starting point, as in vivo assessment of organ/tissue function is a critical aspect for interpretation of the in vitro observations. We have utilized both large (canine) (Bertram et al., 2008,

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14.4 Methods of studying tissue-engineered bladder constructs. A scaffold is seeded with urothelial cells and smooth muscle cells, and implanted into an animal model (center). (a) After the animal is killed, the bladder can be assessed at the tissue level: histologically by slicing and staining the tissue (left), functionally by cutting the tissue into strips and measuring the response to electrical or pharmacological stimuli (middle), and at the molecular level by evaluating gene and protein expression of regenerated tissue (right). (b) In vivo, the bladder can be assessed at the whole-organ level: morphologically with CT imaging (left), and functionally with in vivo urodynamic studies (right). (c) The tissue-engineered bladder can also be assessed at the cellular level by analyzing cell communication via dual whole cell patch clamp methods (left), cell signaling via calcium imaging studies (middle), and by single whole cell patch clamp methods as well (right). Cell photographs courtesy of Dr Yuanyuan Zhang. Bottom right panel modified from Venkateswarlu et al. (2002).

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Oberpenning et al., 1999) and small (rodent) (Burmeister et al., 2008) animal models for this purpose. Our overall research strategy is outlined in Fig. 14.4, which illustrates how in vivo measures are combined with the in vitro techniques described above to provide a more complete picture of the physiological relevance of bladder tissue engineering and regeneration.

14.4.1 Evaluating tissue-engineered bladders For the large animal studies, the canine has been the model of choice. In a seminal study, (Oberpenning et al., 1999) tissue engineering methods were used to perform a subtotal (i.e. trigone sparring) bladder augmentation. A neo-bladder was grown in the laboratory using autologous urothelial and smooth muscle cells seeded onto a polymeric bladder-shaped scaffold. Functional evaluations in vivo (urodynamic studies) for up to 11 months post-implantation of the engineered bladder revealed that the neo-organs/ bladders had an apparently normal capacity to retain urine, normal elastic properties and histological architecture comparable with that of a native bladder. In a more recent investigation Bertram and colleagues (2008) have documented that following radical cystectomy and implantation of a neobladder construct, canines were clinically healthy, continent and able to urinate by 21 days after implantation. Moreover, by 51 days postimplantation, the regenerated bladder wall had histological characteristics consistent with native bladder, including mucosal and serosal lining, detrusor muscle, vascular and nerve composition. Logistic analysis of pharmacological data collected on harvested detrusor smooth muscle strips at 51, 89 and 179 days post-implantation revealed that the sensitivity (EC50 or the concentration of agonist that produces 50% of the calculated maximal response) of contractile responses to carbachol (a non-hydrolyzable acetylcholine analog) and EFS (stimulation of neurotransmitter release from intramural nerve terminals in the bladder wall) was similar to that observed in native bladder in all cases. However, a statistically significant and progressive increase in the calculated maximal response (EMax) to both carbachol and EFS was observed over time. More specifically, although contractile responses to both carbachol and EFS at 51 and 89 days post-implantation were lower in engineered neo-bladder tissues than in native bladder; carbachol-induced contractions, but not EFS-induced contractions, were equivalent to those in native tissue at 179 days post-implantation. Taken together, these data clearly illustrate that the major signaling mechanism responsible for bladder contraction in vivo (i.e. muscarinic receptormediated) is intact in the regenerating bladder in vivo. In short, while further documentation is clearly required, these data point toward the utility of using such an approach. At present, the data are consistent with the supposition that a tissue-engineered neo-bladder is capable of develop-

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ing into a urinary bladder with structural and pharmacological features similar to native bladder within 6 months after surgical implantation.

14.4.2 Evaluating de novo bladder regeneration The aforementioned study describes regeneration/integration of an engineered tissue/organ with the host in vivo, but what about de novo tissue regeneration? In order to begin to evaluate this process, we have taken a similar approach. In this instance, our model of choice is the rodent. Specifically, in the rodent model we are tracking bladder regeneration following a surgical procedure that removes a large portion of the bladder (Burmeister et al., 2008). That is, the entire bladder is excised with the exception of the trigone; for the purposes of this report we will refer to this procedure as a subtotal (trigone-sparing) cystectomy or STC. In short, our preliminary observations with immunohistochemical methods have revealed a population of c-kit (CD117)-positive cells in the detrusor smooth muscle of the bladder wall following STC. While further experiments are clearly required, these c-kit-positive cells presumably reflect the involvement of stem cells in the proliferative portion of the bladder regeneration response to injury. That is, the increased capacity/size of the bladder is not merely due to stretching, but rather involves some proliferative regenerative component. Further characterization of these cells with other antibodies will elucidate more precisely the identity of these cells. Again, the overall goal of this approach is to understand the mechanistic basis for bladder regeneration de novo and use this information to better harness the body’s natural ability to generate/regenerate organs. As described above and depicted in Fig. 14.2, functional studies with tissue derived from regenerating tissue in animal models are also possible by excising the bladder, and suspending detrusor strips in a physiological solution at 37 °C in an organ bath system. The contractility of the muscle in response to pharmacological agents or EFS can be assessed and compared with detrusor contractility from native bladder that was removed from the same animal. Initial observations in the STC rodent model indicate that regenerated detrusor strips exhibit diminished contractility in response to both cholinergic and electrical stimuli 2 weeks following STC (Burmeister et al., 2008). The contractility of the smooth muscle strips improves somewhat up to 8 weeks after the surgery, but it is not yet known whether the contractile response of detrusor smooth muscle from regenerated bladders is able to achieve the level of contractility observed in that same animal prior to STC (i.e. native bladder tissue from the same animal). While these in vitro methods provide valuable insight into the signaling mechanisms and function of the cells that make up the bladder wall, they are not entirely translatable to the clinic, and point toward the importance

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of conducting parallel functional studies in vivo. As with the canine model, urodynamic studies in rodents can also provide valuable insight into the functional characteristics of engineered or regenerating tissue. Studies of bladder function in vivo are widely used in rodents. More specifically, cystometry (i.e. urodynamic studies) on conscious and freely moving rats provides ten distinct parameters that can be used to evaluate bladder function/ performance; both volume and pressure measurements are provided (Boczko et al., 2005, Christ et al., 2001, Christ et al., 2003, Christ et al., 2006, Melman et al., 2005, Pandita et al., 2000, Schroder et al., 2003, Schroder et al., 2004, Woodman et al., 2004, Zotova et al., 2007). In this regard, cystometric analysis has revealed that bladders regenerated in vivo following STC generate micturition pressures that are significantly lower than micturition pressures generated by native bladders. This observation is consistent with the reduced contractility of detrusor muscle strips observed on these same animals in vitro. Interestingly, despite the reduction in pressure generation, the bladders are still able to empty completely. As noted above, the use of a small animal model, such as the rodent also opens up opportunities to examine the utility of non-invasive measures to further augment our understanding of tissue regeneration. In the STC rodent model of in vivo bladder regeneration, for example, micro-computed tomography scans have shown that the bladder does regain a size that approaches, but does not exceed, original values in that same animal. However, also of interest is that the shape of the regenerating bladder appears distinct from that of the native (original) bladder in that same animal, in that it is more spherical than pear-shaped. The implications of this distinct shape for long-term bladder function remain to be elucidated. Consistent with the measurements made using traditional urodynamic studies micro-computed tomography also revealed a positive correlation between micturition pressure and bladder circumference as measured by the anterior view (Burmeister et al., 2008). While these preliminary data are clearly intriguing and further work is underway, our observations are certainly consistent with the possibility that non-invasive measures of bladder regeneration may provide important indices (i.e. metrics) with which to track the progress of the bladder regeneration process, as well as provide mechanistic insight into the process. In addition to CT scans, we anticipate that utilization of magnetic resonance imaging (MRI) methods will also provide valuable insight into how bladder regeneration progresses, and the subsequent integration with the host tissue.

14.5

Conclusions and future trends

The rate of application of tissue engineering/regenerative medicine technologies to urological indications continues to accelerate. The parallel

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development of enabling technologies (i.e. scaffolds, bioreactors, improved drug delivery methods, nanotechnology and non-invasive measures of tissue regeneration) is expected to further accelerate progress. The use of a wellestablished algorithm to assess objectively and enhance the performance of the engineered and regenerating tissues should ensure an iterative improvement process with the ultimate end point being ‘biomimetic’ bladders that closely approximate native bladder with respect to structure, function and longevity.

14.6

References

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15 Regenerative pharmacology and bladder regeneration K.- E. A N D E R S S O N and G. J. C H R I S T, Wake Forest Institute for Regenerative Medicine, USA

Abstract: Pharmacological sciences are currently applied to regenerative medicine/tissue engineering on a goal-directed, tissue-specific basis. However, regenerative pharmacology has a wider scope and can be broadly defined as the application of pharmacological sciences to the many different areas of regenerative medicine/tissue engineering – such as cell and molecular biology, molecular genetics, chemical and material sciences, engineering and transplantation. The ultimate aim of regenerative pharmacology is to modulate cells and tissues in order to improve or enhance their function, and/or accelerate their development. This aim can be achieved by application of either ‘active’ or ‘passive’ approaches, using established pharmacological methods in a systematic way to characterize cells and tissues, endogenously regenerated or bioengineered, for accumulation of new knowledge about receptors, signaling pathways and cell/tissue/organ function. This is illustrated in this chapter by examples from studies of bladder regeneration and tissue engineering of the bladder. There is obviously enormous potential and demand for pharmacology and the pharmacological sciences to shape and accelerate the development of regenerative medicine. Key words: regenerative pharmacology, tissue engineering, bladder regeneration, engineered bladder, neobladder.

15.1

Introduction

15.1.1 Regenerative medicine Regenerative medicine is now a well-established multidisciplinary field, encompassing various areas of technology development, such as tissue engineering, stem cells and cloning. Tissue engineering is one of the major components of regenerative medicine and follows the principles of cell transplantation, materials science and engineering towards the development of biological substitutes that can restore and maintain normal function; it includes enhancement of endogenous regenerative capacity (Atala, 2007; Yamzon et al., 2008). Current themes in the field include the use of 322

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cells seeded on to biocompatible absorbable matrices for implantation as tissue substitutes, which is conducive to host tissue ingrowth. Injection therapy of cells for organ rehabilitation is also making strong headway towards the restoration of organ structure and function. In regenerative medicine and the tissue engineering process, pharmacological principles and tools are commonly used, but their full potential has not yet been fully exploited/realized.

15.1.2 What is regenerative pharmacology? Regenerative pharmacology can be broadly defined as the application of pharmacological sciences to the different areas of regenerative medicine, including cell and molecular biology, molecular genetics, chemical and material sciences, engineering and transplantation. The ultimate aim is to modulate cells and tissues pharmacologically to improve or enhance their function and/or accelerate their development. Generally, two approaches may be used: (a) the ‘active’ approach – use of growth factors and different pharmacological agents to alter (i.e. enhance or depress), cell growth, differentiation and function; and (b) the ‘passive’ approach – using established pharmacological methods to characterize cells and tissues, endogenously regenerated or bioengineered. Both of these approaches are currently used in regenerative medicine; however, what is needed is application of these approaches in a systematic way to accumulate new knowledge about receptors, signaling pathways, and cell/tissue/organ function. This is a new and exciting area with a great potential to improve and enhance our understanding of the bioengineering/regenerative process, as well as the quality of bioengineered/regenerated tissues. The use of stem cells from different sources has shed new light on the understanding of mechanisms responsible for various disorders, and has opened a new field for regenerative medicine and regenerative pharmacology. The mechanisms that govern stem cell fate decisions are under tight control but remain potentially alterable. Recent studies have shown that several currently used drugs such as colony stimulating factors, statins, angiotensin-II receptor antagonists/angiotensin converting enzyme (ACE) inhibitors, erythropoietin, nitric oxide donors, estrogens and glitazones, have modulatory activity on stem cell functions (Romagnani et al., 2007). These drugs mostly enhance stem cell survival and mobilization. Furthermore, a series of new pharmacological agents – such as the chemokine receptor antagonist AMD3100, glycogen synthase kinase-3 (GSK-3) inhibitors and histone deacetylase inhibitors (HDACi), which modulate the growth, differentiation and mobilization of stem cells – are currently under evaluation in both in vivo experimental models and preliminary clinical trials (Romagnani et al., 2007).

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The application of pharmacological principles to accelerate and optimize (either in vitro or in vivo) the maturation and function of bioengineered tissues was recently discussed (Andersson and Christ, 2007). As such, pharmacology should assume a critical role in all aspects of the regenerative process (Fig. 15.1). As noted above, pharmacological methods can facilitate the development of the enabling technologies required for accelerated cellular and tissue maturation and function, as well as characterize bioengineered tissue function to ensure their reliability and utility. Promising laboratory discoveries have exciting implications for novel clinical applications of regenerative medicine in the genitourinary tract (Atala, 2007; Yamzon et al., 2008). As an example of the application of pharmacological methods to studies of bladder regeneration as well as tissue engineering, we will briefly review some recent developments in this field, illustrating how pharmacological tools and methods can be iteratively applied. The focus here is on bioengineered constructs in which the smooth muscle cell is a primary parenchymal cell type and, thus, smooth muscle cell tone (i.e. contraction and relaxation) is of major importance. However, the principles outlined here will have broad applications for regenerative medicine that may involve other cell types and engineered tissues.

15.2

Endogenous bladder regeneration

The possibility to enhance the endogenous regenerative capacity of an organ such as the bladder, and to characterize pharmacologically and functionally the newly synthesized tissue offers a challenge for regenerative pharmacology. Studies to ‘actively’ enhance endogenous bladder regeneration still seem to be lacking, but this is not surprising given the paucity of information about the mechanistic basis of de novo bladder regeneration in vivo. Nonetheless, investigations using a ‘passive’ approach are emerging. For example, Frederiksen et al. (2004) studied the pharmacological and mechanical properties of newly developed detrusor muscle after subtotal cystectomy. They performed a partial, trigone-sparing cystectomy in female rats and explored, by pharmacological means, whether the regenerated detrusor had characteristics similar to the normal bladder base (supratrigonal segment), from which it regenerated, or to the normal bladder body (equatorial segment), which it replaced (see Fig. 15.2). Fifteen weeks after the operation, detrusor strips were cut from supratrigonal and equatorial segments (middle of the bladder body). Shamoperated rats served as controls. Responses to electrical field stimulation (EFS) were obtained in the absence of scopolamine (blocking muscarinic receptors), prazosin (blocking α1-adreneceptors) and α–β-methylene-ATP (desensitizing P2X1 receptors). Concentration–response curves were obtained for carbachol (stimulating muscarinic receptors), α–β-methylene-

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Regenerative medicine

Cell expansion and differentiation 1

2

Biomaterials 3

Regenerative pharmacology

Tissue biopsy

Enabling technologies 5

Implantation

4

Clinical studies

15.1 The central role of regenerative pharmacology in regenerative medicine and tissue engineering. 1 ‘Cell expansion and differentiation’ signifies cell expansion and differentiation via modulation/ manipulation of growth factors, cytokines, hormones, neurotransmitters and neuromodulators – both before and after scaffold seeding. 2 ‘Biomaterials’ involve the use of biodegradable, biocompatible, cell- and tissue-specific scaffolds to modulate cellular phenotype/pharmacology and promote cellular integration and tissue formation in vitro and in vivo. 3 ‘Enabling technologies’ will use nanotechnology in the development of ‘smart’ scaffolds for sustained or burst release of drugs, hormones, chemicals and growth factors. Bioreactors are center-stage for tissue preconditioning and optimization of cellular phenotype, tissue organization, maturation and function, as is the development of non-invasive pharmacological evaluation of engineered tissues. Thus, it is clear that enabling technologies will be required throughout the tissue engineering process, and are merely illustrated at this point in the process to emphasize the importance of this particular step to mainstream pharmacology (see Table 15.1). 4 ‘Implantation’ includes the functional evaluation of engineered tissues. Again, drug and cell delivery technologies can be utilized to further accelerate and enhance tissue maturation, integration and function. Vascular endothelial growth factor (VEGF) and nerve growth factor (NGF), for example, may be provided via encapsulation techniques to promote accelerated growth of vessels and nerves, respectively, into the engineered tissue. 5 ‘Tissue biopsy involves the pharmacological characterization of cell phenotype and tissue function. Again, during the preclinical development phase of the tissue engineering process, one would assess membrane receptors, signal transduction and excitation– contraction coupling. This information is then used to evaluate the existing protocol and thereby direct optimization of the tissue engineering process.

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Supratrigonal

Equatorial

15.2 There is a regional variation in the bladder in response to electrical field stimulation and to drugs stimulating muscarinic, α-adrenergic and purinergic receptors.

ATP (stimulating P2X1 receptors) and phenylephrine (stimulating α1-adrenoceptors). The results revealed that the maximal contractile response to EFS was 60% of that to high-K+ solution (i.e. depolarization with KCl) in strips from both control and cystectomy bladders. Prazosin had no effect. Scopolamine decreased the maximal response of supratrigonal strips to 62% (controls or sham-operated) and 61% (operated or subtotal cystectomy) of that without blocker. For equatorial strips the decrease was to 81% (controls or shamoperated) and 58% (operated or subtotal cystectomy). Frequency–response relations were obtained during blockade with scopolamine, α–β-methyleneATP and prazosin. Supratrigonal strips showed a pronounced additional inhibition up to 40 Hz. Equatorial strips from controls were completely inhibited at all frequencies. Equatorial strips from operated bladders were inhibited up to 20 Hz but not at 40 and 60 Hz. Computer analysis of the concentration-response curves, using the logistic equation, revealed that the carbachol EC50 values (concentration of carbachol that produced a response equivalent to 50% of the calculated maximal response) were similar in all groups. Moreover the calculated maximum response to phenylephrine (Emax) was 10–20% of the high-K+ response. The authors concluded that there was a regional difference in pharmacological properties of normal detrusor, with a considerable contractile response to stimulation remaining in the supratrigonal muscle after simultaneous cholinergic, adrenergic and purinergic blockade. The new detrusor seemed functionally well innervated with no supersensitivity to muscarinic stimulation, and the newly formed

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bladder body had pharmacological properties specific for the supratrigonal segment from which it had developed. In another initial report, urodynamic studies to evaluate bladder function in vivo, revealed that regenerated bladders were able to empty completely within 8 weeks after a subtotal cystectomy, albeit with slightly reduced volumes and with reduced maximal intravesical pressure (Burmeister et al., 2008; See Chapter 14 in this volume for more details). Taken together, the data suggest that this model is attractive for in vivo studies aiming to elucidate the mechanistic basis for de novo bladder regeneration in vivo. Such studies are an absolute prerequisite for developing pharmacological methods and technologies that can stimulate and enhance endogenous regeneration of bladder tissue by local or systemic administration of, for example, growth factors, cytokines, hormones, neurotransmitters and neuromodulators.

15.3

Construction of a tissue or organ

The construction of bioengineered tissues is a multi-step process (Atala 2006; Mikos et al., 2006). As illustrated in Fig. 15.1, the first step is the identification and harvest of somatic cells, for which autologous (and thus immunocompatible) cells derived from the tissue of interest would be the preferred choice. However, there will inevitably be instances in which biopsies are not possible or in which tissues are so badly damaged that isolation and expansion of autologous cells from those tissues becomes impossible. In such situations, stem cells, derived from embryonic, amniotic, cord blood, placental or bone marrow sources, become attractive alternatives (Atala 2006; Mikos et al., 2006; De Coppi et al., 2007). The possibility, using pharmacological means (Romagnani et al., 2007), of enhancing stem cell survival and mobilization and modulating their growth and differentiation, opens new approaches to the next steps in the process involving cell growth and expansion. The goal is to derive sufficient numbers of cells with the appropriate phenotypic and functional characteristics that can be ‘seeded’ on to a matrix system, or scaffold. The cell-seeded scaffolds are then either directly implanted into the host, or they may be further incubated under physiological conditions; in the latter case, cell-seeded scaffolds can be preconditioned in ‘bioreactors’. Regardless of the precise incubation conditions, the concept remains the same – once the bioengineered constructs have been appropriately acclimated or matured, they are then implanted into the host. In any event the challenge for regenerative pharmacology is to control/modulate cell growth, expansion and differentiation by using, for example, growth factors, cytokines, hormones, neurotransmitters and neuromodulators, and subsequently to evaluate the cellular phenotype and function pharmacologically before implantation.

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15.4

Development of an engineered bladder

Tissue-engineered neobladders were successfully generated and implanted in dogs undergoing trigone-sparing cystectomy, exhibiting full functional capacity (Oberpenning et al., 1999). Such bladders have also been successfully implanted in patients, with up to 5 years of follow-up, supporting the potential of the urological applications of regenerative medicine (Atala et al., 2006). These studies followed the process outlined in Fig. 15.1 and represented the culmination of more than a decade of preclinical work.

15.4.1 Cellular source Loss of some salient phenotypic characteristics when placed into cell culture environments is a common finding for smooth muscle cells (Owens 1995; Owens et al., 2004), including bladder smooth muscle cells (Kropp et al., 1999). For example, when maintained in two-dimensional cultures, bladder myocytes dedifferentiate and lose their ability to contract in response to appropriate agonists (Kropp et al., 1999). In addition, it is not known to what extent dedifferentiated smooth muscle cells are able to regain their normal phenotypic characteristics after implantation into the host. This is a major concern given that cellular phenotype will determine proper functioning of engineered bladder tissue after implantation. In this regard, Kropp et al. (1999) characterized the in vitro contractile response of human and rat bladder smooth muscle cells to several pharmacological agonists known to induce in vivo contraction of intact bladder muscle. To this end, human and rat bladder smooth muscle cells were seeded separately within attached collagen lattices. Subsequently, the contractile ability of these cells was evaluated by evaluating alterations in lattice diameter after exposure to a variety of relevant modulators of detrusor contractility – namely, carbachol, calcium ionophore, lysophosphatidic acid (LPA), endothelin, KC1, angiotensin and serotonin. Results were recorded as a mean reduction of the lattice diameter. In short, human smooth muscle cells were contracted by calcium ionophore, LPA and endothelin, but not by carbachol, angiotensin, KC1 or serotonin. Rat bladder smooth muscle cells had a similar contractile response but did not contract in response to endothelin. Kropp et al. (1999) concluded that in vitro cultured bladder smooth muscle cells demonstrated a loss of contractile response to normal in vivo pharmacological agonists. Their results demonstrated the ability to further characterize cultured muscle cells by in vitro contractility studies, and they concluded that such characterization is essential for advancing our understanding of the clinical applicability of in vitro studies utilizing cultured bladder smooth muscle cells. The obvious challenge for regenerative pharmacology is to find ways

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to maintain the contractile phenotype, with respect to receptor-mediated signal transduction, as well as the intracellular signaling pathways that they activate. The phenotype of the cultured smooth muscle cell should be quite well established before implantation, if one wishes to evaluate the impact of the body’s ‘bioreactor’ on the tissue engineering process. In the end, improvements in the tissue engineering process as well as development of requisite enabling technologies, will depend on more detailed knowledge of how closely in vitro engineered tissues approximate the native tissues that they will replace. Lai and coworkers (2002) studied human smooth muscle cells derived from functionally normal or dysfunctional bladders (i.e. exstrophic bladders and neurogenic bladders). The goal of this investigation was to evaluate the suitability of these distinct cell sources for bladder tissue engineering. These researchers cultivated, expanded and seeded bladder cells from both patient populations on to biodegradable polymer scaffolds, which were then implanted into athymic nude mice. After 2 months, the cell-seeded scaffolds demonstrated contraction and relaxation responses to supramaximal electrical field and carbachol stimulation that were reminiscent of the expected native responses. The authors concluded that smooth muscle cells from normal and diseased bladders retain important aspects of their phenotype in vitro as well as after implantation. That is, that tissue engineered constructs derived from either urodynamically normal or pathological bladders yielded essentially equivalent physiological/pharmacological results. In short, these observations suggest that bladder muscle cells, regardless of their origin, may have the potential to be engineered into functional bladder tissues.

15.4.2 Biomaterials and scaffolds Tissue engineering generally requires an artificial extracellular matrix (ECM) to promote the tissue tissue regeneration process. Moreover, this artificial ECM should disappear through degradation/absorption into the body when the new tissue is regenerated (i.e. remodeling). Thus, it is important that the materials used for the matrix should be prepared from biodegradable polymers. This requirement, as well as the need for adequate cell adhesion on to the matrix surface, makes biological materials attractive in tissue engineering. The development and application of biomaterials are thus essential to tissue engineering in regenerative medicine (Atala and Lanza, 2002; Rosso et al., 2005; Stayton et al., 2005; Furth and Atala, 2007). In addition, scaffolds provide mechanical support and shape for the developing bioengineered tissue in vitro and post-implantation, allowing seeded cells to expand, differentiate and organize. In this regard, three main types of biomaterials have been used for tissue engineering: (a) naturally derived

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materials (e.g. fibrin, collagen, chitosan and alginate); (b) synthetic polymers (e.g. polyglycolic acid (PGA), polylactic acid (PLA), polylactic-coglycolic acid (PLGA), polyhydroxyalkanoate (PHA) and polydioxanone (PDS); and (c) acellular tissue matrices (e.g. bladder lamina propria and small intestinal submucosa). In the engineering of neobladders, for example, synthetic polymers have been utilized as scaffolds. This is an important concept, as it permits the macrostructure, mechanical properties and degradation (i.e. remodelling) process to be readily controlled and manipulated. In this regard, Oberpenning and colleagues (1999), used a novel bladder augmentation technique to reconstitute the first functional dog urinary bladder. For these investigations, they used PGA matrix configured into a bladdershaped mold, which was coated with PLGA in order to achieve adequate mechanical strength. A biodegradable composite scaffold made of collagen and PGA/PLGA was also used for implantation of bladder constructs in humans by Atala and coworkers (2006). The true importance of such biodegradable scaffolds is that their degradation can prevent a chronic foreign body response.

15.4.3 Development of ‘smart’ scaffolds Further optimization of the regenerative process requires the development of several critical enabling technologies (see Fig. 15.1). Among these, there is a strong need for drug delivery systems that can leverage available nanotechnology to provide biomaterials and tissue engineering scaffolds that contain ‘biological signals’ (nerve and vascular growth factors, cytokines for attracting endogenous stem cells, etc.) that can further augment/accelerate the tissue engineering/regenerative medicine process. This will require novel systems that can efficiently and precisely deliver relevant biomolecules to intracellular targets. Several such approaches have been described for the development of such ‘smart’ materials and scaffolds (Rosso et al., 2005; Stayton et al. 2005; Furth and Atala 2007; Moroni et al., 2008). As mentioned above, nanotechnology seems to be a particularly powerful tool for creating these materials and scaffolds (Engel et al., 2008; see Chapter 13 in this volume). In short, the potential of integrating, for example, growth factors and pharmacological agents into the scaffolds to promote and enhance cell growth and differentiation is an obvious challenge for regenerative pharmacology and is currently under exploration.

15.5

Implantation of the bladder construct in preclinical studies

After configuration of the bladder mold, the scaffold is typically seeded with urothelial cells on the inside of the scaffold and smooth muscle

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cells on the outside of the scaffold; it is then placed in an incubator for a period of 3–4 days. At that point, the bladder construct is largely coated with cells and can be implanted in the host (i.e. dog or human (see below)). In functional evaluations for up to 11 months after implantation in the dog, the neobladders demonstrated a normal capacity to retain urine, normal elastic properties and normal histological architecture (Oberpenning et al., 1999). With respect to the preclinical studies, the application of integrative and organ systems pharmacology and the use of imaging and functional assessment modalities, offers exciting possibilities for functional evaluation and monitoring of the engineered tissues. The application of non-invasive imaging techniques to the tissue engineering/regenerative medicine process is still in its infancy, but is clearly expected to provide important metrics for both predicting and improving the outcomes. Moreover, it should also be possible to further optimize tissue/organ function in vivo by using agents known to increase, for example, bladder contractility, when required and appropriate.

15.6

Preliminary clinical experience with neobladders

In patients with end-stage bladder disease, gastrointestinal tissue can be used for bladder reconstruction, but the complications associated with this process of cystoplasty can be numerous. As an alternative process, Atala and colleagues (2006) have successfully used autologous engineered bladder tissues for the reconstruction of bladders in seven patients (aged 4–19 years, all with myelomeningocele) with high-pressure or poorly compliant bladders. The investigators used bladder biopsies from each patient to establish urothelial and smooth muscle cells in culture; cells were then seeded on a biodegradable bladder-shaped collagenous scaffold. About 7 weeks after initial biopsy, the autologous engineered bladder constructs were used for reconstruction and implanted either with or without an omental wrap. One to five years after implantation, the engineered bladders with an omental wrap manifested superior characteristics (bladder leak point pressure decrease at capacity, volume increase and compliance increase) relative to those without omental wrap. Consistent with the observed improvements in bladder volume and compliance, biopsies of the engineered bladders revealed adequate structural architecture and phenotype (Atala et al., 2006). In addition, a phase-II clinical trial of neobladders engineered according to this process has been initiated under an investigational new drug (IND) petition approved by the Food and Drug Administration (FDA) (www.Tengion.com; www. clinicaltrials.gov). These studies are being conducted in patients with neurogenic bladders secondary to spina bifida (myelodysplasia).

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15.7

Conclusions

To date, pharmacological sciences have been applied to regenerative medicine/tissue engineering on a goal-directed, tissue-specific basis. However, what is required to achieve the ultimate potential of these groundbreaking regenerative technologies is a more systematic application of pharmacological principles at each of the different steps in the process. When complimented with the host of cutting-edge multidisciplinary technologies already available in this important area of medical research, there is obviously enormous potential and demand for pharmacology and the pharmacological sciences to shape and accelerate the development of regenerative medicine.

15.8

Acknowledgement

We would like to thank Dr James Yoo for valuable comments and suggestions.

15.9

References

andersson ke and christ gj (2007), Regenerative pharmacology: the future is now. Mol Interv Apr;7(2):79–86. atala a (2006), Recent applications of regenerative medicine to urologic structures and related tissues. Curr Opin Urol Jul;16(4):305–9. atala a (2007), Engineering tissues, organs and cells. J Tissue Eng Regen Med Mar–Apr;1(2):83–96. atala a and lanza rp (2002) Methods of Tissue Engineering. Academic Press, New York. atala a, bauer sb, soker s, yoo jj and retik ab (2006), Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet Apr 15;367(9518):1241–6. burmeister d, aboushwareb t, andersson k-e and christ g (2008), Characterization of normal rat bladder regeneration in an in vivo model. Presented at the Experimental Biology Meeting, San Diego, CA, April 5–9, poster B26 581.2. de coppi p, bartsch g jr, siddiqui mm, xu t, santos cc, perin l, mostoslavsky g, serre ac, snyder ey, yoo jj, furth me, soker s and atala a (2007), Isolation of amniotic stem cell lines with potential for therapy. Nat Biotechnol Jan;25(1):100–6. engel e, michiardi a, navarro m, lacroix d and planell ja (2008), Nanotechnology in regenerative medicine: the materials side. Trends Biotechnol Jan;26(1):39–47. frederiksen h, arner a, malmquist u, scott rs and uvelius b (2004), Nerve induced responses and force-velocity relations of regenerated detrusor muscle after subtotal cystectomy in the rat. Neurourol Urodyn 23(2):159–65. furth m and atala a (2007), Tissue engineering: future perspectives. In Principles of Tissue Engineering, eds Lanz R, Langer R and Vacanti J, 3rd edition, pp. 33–50, Elsevier Academic Press, Oxford.

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kropp bp, zhang y, tomasek jj, cowan r, furness pd 3rd, vaughan mb, parizi m and cheng ey (1999), Characterization of cultured bladder smooth muscle cells: assessment of in vitro contractility. J Urol Nov;162(5):1779–84. lai jy, yoon cy, yoo jj, wulf t and atala a (2002), Phenotypic and functional characterization of in vivo tissue engineered smooth muscle from normal and pathological bladders. J Urol Oct;168(4 Pt 2):1853–7. mikos ag, herring sw, ochareon p, elisseeff j, lu hh, kandel r, schoen fj, toner m, mooney d, atala a, van dyke me, kaplan d and vunjak-novakovic g (2006), Engineering complex tissues. Tissue Eng Dec;12(12):3307–39. moroni l, de wijn jr and van blitterswijk ca (2008), Integrating novel technologies to fabricate smart scaffolds. J Biomater Sci Polym Edn 19(5):543–72. oberpenning f, meng j, yoo jj and atala a (1999), De novo reconstitution of a functional mammalian urinary bladder by tissue engineering. Nat Biotechnol Feb;17(2):149–55. owens gk (1995), Regulation of differentiation of vascular smooth muscle cells. Physiol Rev Jul;75(3):487–517. owens gk, kumar ms and wamhoff br (2004), Molecular regulation of vascular smooth muscle cell differentiation in development and disease. Physiol Rev Jul;84(3):767–801. romagnani p, lasagni l, mazzinghi b, lazzeri e and romagnani s (2007), Pharmacological modulation of stem cell function. Curr Med Chem 14(10):1129–39. rosso f, marino g, giordano a, barbarisi m, parmeggiani d and barbarisi a (2005), Smart materials as scaffolds for tissue engineering. J Cell Physiol Jun;203(3):465–70. stayton ps, el-sayed me, murthy n, bulmus v, lackey c, cheung c and hoffman as (2005), ‘Smart’ delivery systems for biomolecular therapeutics. Orthod Craniofac Res Aug;8(3):219–25. yamzon jl, kokorowski p and koh cj (2008), Stem cells and tissue engineering applications of the genitourinary tract. Pediatr Res May;63(5):472–477.

16 Autologous cell sources for urological applications Y. Z H A N G, Wake Forest Institute for Regenerative Medicine, USA

Abstract: Recent advances in regenerative medicine, including novel approaches to urological tissue engineering, have accelerated the process of translation of cell-based therapy from the bench to the clinic. This chapter reviews the progress and challenges of in vitro tissue engineering and in vivo tissue regeneration with autologous cells in urological applications. Use of adult stem cells from patients’ own tissue for regeneration is rapidly becoming a promising technology because it lacks many of the drawbacks of non-autologous tissue replacement techniques, such as immune rejection and ethical or religious concerns. These cells are obtained from various tissue types, and include urological stem and/or progenitor cells as well as cells from non-urological tissues. Methods of directing stem cells to differentiate into specialized cell types, delivering factors that stimulate cell proliferation, remodeling urological tissues and synthesizing new extracellular matrix are under investigation. Several populations of autologous differentiated cells, adult stem cells and progenitor cells have been investigated as potential cellular sources in the reconstruction of urological tissues and organs. Nevertheless, considerable basic research is still required to ensure the specific differentiation, long-term fate and functionality of stem cells after transplantation. However, based on recent promising reports from the clinic, autologous cell-based therapy could soon become a standard procedure in the treatment of urological diseases. Key words: stem cells, biomaterials, tissue engineering, cell therapy, urological diseases.

16.1

Introduction

Significant progress has been made in the field of urological tissue engineering and cell therapy over the past two decades. Urological tissue regeneration strategies include several approaches: cell-seeded and cell-free tissue engineering, cell therapy, and stem cells and cloning technology.1,2 Cellbased tissue engineering uses scaffolds seeded in vitro with primary cells obtained from a tissue biopsy. This composite graft is then implanted back into the host for carry-over of the regenerative process.3–8 It has been shown 334

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that cell-seeded tissue engineering may be beneficial for large-size replacement of bladder tissue and in tubular urethral reconstruction. The nonseeded technology uses cell-free biodegradable scaffolds that allow the natural process of regeneration to occur in vivo, and it has been successfully applied experimentally and clinically in partial or small-size tissue replacement for urethral9 and bladder repair.10,11 Recently, clinical trials of stem/ progenitor cell therapy techniques for bladder reconstruction, urinary stress incontinence (USI) and vesicoureteral reflux (VUR) have produced encouraging results. These tissue engineering approaches could allow dysfunctional tissues to be repaired or replaced in patients with various urological congenital defects, traumatic injuries, and chronic infections, as well as in those patients who require reconstructive procedures after cancer surgeries. In cell-based tissue engineering, the grafted cells can be from an individual’s own tissue (autologous), from the same species but a different individual (allogenic), or from a different species (xenogenic). The cell sources – which can include embryonic, fetal and adult stem cells as well as fully differentiated cells – are currently being investigated for tissue engineering use. Embryonic stem cells are thought to have the most capability because of their pluripotent properties12–14 – they can theoretically grow into any of over 200 cell types from all three germ layers. In addition, embryonic stem cells are self-renewing. Although human embryonic stem cells are preferred for tissue engineering purposes, questions of immunecompatibility, possible tumor (teratoma) formation in vivo, and ethics remain. Thus, other cell types have been derived from near-mature tissues, including amniotic fluid, placenta, and umbilical cord blood. These cells provide alternative cell sources for autologous therapies for the donor, as well as for individuals who match the tissue types of these stem cells. However, although these allogenic and xenogenic cells are easier to obtain, use of these cells may still entail risky immune-suppression regimens. An ideal cell source for use in engineered urological constructs should be non-immunogenic, easy to harvest, proliferative, and capable of differentiating into functional, mature cells. Several cell sources have been proposed for tissue engineering. Autologous cells, including mature and stem/progenitor cells, have several advantages. The use of these cells avoids rejection and thus avoids the deleterious complications of immunesuppressive medications, and the risk of infectious disease transmission is eliminated. Therefore, autologous cells are considered as a ‘gold standard’ source for urological tissue regeneration.6,7,15–24 Recently, there have been increasing reports in the literature about the use of autologous stem/ progenitor cells in urological tissue regeneration, i.e. urine-derived progenitor cells,25 and non-urological stem cells including adipose-derived stem cells,26–28 bone marrow-derived stem cells (BMSC),6,29 skin-derived progenitor cells,30 and skeletal muscle-derived progenitor cells31,32 (Table 16.1).

Stem/progenitor cells

Fully differentiated cells Yes Yes Yes

Buccal mucosa Skin

Urothelium from bladder wash Chrondocytes Blood cells Yes

Yes

Yes

No

Skeletal muscle-derived progenitor cells

Bone marrow stem cells

Adipose stem cells

Urine progenitor cells

Yes No

Yes

Require tissue biopsy?

Bladder cells

Cell types

Yes

Depends on age

No

Limitation of growth in vitro

Yes

Highly expandable?

Cell therapy for VRU and urinary incontinent bladder regeneration Tissue-engineered bladder; cell therapy for VRU; urinary incontinence Tissue-engineered; bladder; VRU, source of smooth muscle cells Tissue-engineered urological organs

Cell therapy for VRU Cell therapy for VRU

Urethral repair

Tissue-engineered urological organs Hypospadias or urethra repair Hypospadias or urethra repair

Potential applications

Table 16.1 Potential use of autologous cells for urological tissue engineering and cell therapy

No

Yes

Yes

Yes

Yes Yes, unsuccessful

No

Yes Yes

Yes

Clinical trial?

Autologous cell sources for urological applications

16.2

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Fully differentiated cells for urological reconstruction

16.2.1 Cells derived from urological tissues In cell-based urological tissue engineering, autologous urological cells, i.e. urothelial and smooth muscle cells, are the primary cell sources currently in use. Urothelial cells Urothelial cells compose a specialized epithelium that plays important and conflicting roles: the urothelium must act as a permeability barrier to protect underlying tissues against noxious urine components while also stretching to accommodate urine pressures. Under normal conditions, urothelial cells turn over and differentiate very slowly. However, if the bladder mucosa is injured, the urothelial cells are able to proliferate and regenerate rapidly and the wound is repaired in a few days. This capacity for regeneration has been observed in vitro. Urothelial cells isolated from normal human urothelium and serially propagated as monolayers in serum-free culture were homogeneous and adopted a proliferative, nondifferentiated phenotype.23,33 In order to establish a successful culture of urological cells, the proper methods of harvesting tissue are critical, although cultivation of urothelium has been well developed.33–35 When a biopsy for obtaining bladder tissues is performed, electrocautery should be avoided since electrocautery injures tissue and may affect the number of cells that survive in the primary culture. Additionally, the presence of bladder foreign bodies, such as catheters or pre-existing stones, can lead to disturbed cell growth in vitro.34 Therefore, a bladder free of foreign bodies is required before tissue biopsy. To provide functional cells for bladder or urethral reconstruction and remodeling, cultured cells at early passage (under passage five) provide the optimal results. Multilayered urothelial sheet Studies have shown that urothelial cells can be induced to form a multilayered urological sheet in the presence of serum or physiological concentrations of calcium.34,35 This tissue-engineered stratified urothelial construct consisted of basal membrane and basal, intermediate, and superficial cell layers. The apical membrane of superficial cells formed villi and glycocalices, and tight junctions and desmosomes were observed. Immunohistochemistry showed similarities and differences in the expression of cytokeratins, integrin, and cellular adhesion proteins. In the cultured

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urothelium, cytokeratin 20 and integrin subunits α6 and β4 were absent, and symplekin was expressed diffusely in all layers. Uroplakins were clearly expressed in the superficial umbrella cells of the urothelial constructs; however, they were also present in intermediate and basal cells. Symplekin and uroplakins were expressed only in the superficial cells of native bladder tissue. The urothelial constructs showed excellent viability, and functionally their permeability for water, urea, and ammonia as no different from values measured in native human urothelium. Proton permeability was even lower in the constructs compared with that of native urothelium. Thus, although the in vitro cultured human stratified urothelium did not show complete terminal differentiation of its superficial cells, it retained the same barrier characteristics against the principal urine components. These results indicate that an in vitro cultured multilayer urothelium sheet, if grown on a compliant degradable support or in co-culture with smooth muscle cells (SMCs), may be suitable for urethral repair or reconstructive cystoplasty. Characterization of urothelial cells Urothelium displays distinctive features associated with its specialized role as a permeability barrier. The progression from basal to superficial cells is accompanied by an ordered succession of phenotypic changes that can be detected with immunocytochemical staining. A group of urothelial umbrella cell-specific markers called ‘uroplakins’ (UPs) consists of four major uroplakin proteins (UP Ia, Ib, II, and III) that form UPIa/UPII and UPIb/UPIII pairs. UPIa/UPII are confined to superficial cells and UPIb alone is expressed by intermediate cells.36 These proteins are synthesized abundantly by normal urothelial cells. Therefore, UPs are the characteristic integral membrane proteins in the terminally differentiated, superficial urothelial asymmetric unit membrane. The other cell markers for urothelial cells are cytokeratin proteins. The cytokeratin family (CK) is generally expressed as a correlate of all epithelial tissue development and is commonly used as a bladder epithelial cell marker to identify urothelial cells. On histological sections of human bladder tissue, monoclonal antibodies to CK 7, 8, 17, 18, and 19 react with all cell layers of bladder mucosa. Antibodies to CK 13 react with the basal and intermediate cells, but not with superficial cells.33 Antibodies to CK 20 react only with superficial cells. Cultured urothelial cells maintain the expression pattern of almost all keratin isotypes present in the normal mucosa tissue. Anti-pancytokeratins, AE1/ AE3s, are better overall cytokeratin markers, and are made up of a cocktail of both high and low molecular weight cytokeratins. AE1 contains CK 10, 13, 14, 15, 16, and 19; and AE3 contains CK 1, 2, 3, 4, 5, 6, 7, and 8. AE1/ AE3 is one of the most commonly used epithelial cell markers for identifing urothelial cells.37,38

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Urological smooth muscle cells SMCs compose a layer of smooth muscle fibers arranged in spiral, longitudinal, and circular bundles in the bladder wall. Cultured bladder SMCs are arranged in sheets or bundles and connected by gap junctions. Most mature cells can only be reproduced for a limited number of cell generations in in vitro tissue culture conditions. Certain cell types, such as SMCs, lose their phenotype at early passages in culture.39 As a result of normal cell aging in tissue culture, the cultured cells eventually lose their specialization in form and/or function (dedifferentiation) or die. In order to maintain the SMC phenotype when creating a tissue-engineered bladder, it is preferable to use SMCs below passage five and culture them on scaffolds for a short period of time (less than 2 weeks). Immortalization of cells establishes continuously proliferating cell lines that play an important role in the study of the biology of cell growth, differentiation, and apoptosis. Immortalized cell lines are usually transformed through the insertion of viral genes, which ultimately causes the cells to grow indefinitely in culture. Although immortalized muscle cell lines could generate a large number of functional cells with muscle marker expression,40 such cell lines have abnormal characteristics and carry a higher risk of tumor formation compared with non-transformed cells. Thus, while these types of cells are useful for cancer research, they have limitations in clinical applications.

Characterization of smooth muscle cells Cytoskeletal and contractile proteins are expressed in differentiated smooth muscle populations and other cell types. These proteins have been commonly used to characterize these cell types and to identify SMC in urological tissues. These SMC markers include alpha-smooth muscle actin (α-SM actin), smoothelin, desmin, myosin heavy chain, calponin, caldesomon, and vimentin. α-SM actin is expressed in SMC in smooth muscle actin and myofibrils that indicate the formation of intracellular ‘stress fibers’. α-SM actin is also expressed in other cell types, such as activated fibroblasts or myofibroblasts in wounds or transplanted hearts, or by proliferating mesangial cells in glomerular disease of the kidney. Thus, α-SM actin protein is not SMCspecific. Although it cannot be used as a definitive marker for the SMC lineage, α-SM actin is still useful as a differentiation marker for SMCs because it is the first cell type-selective protein expressed by SMCs during differentiation, and is one of the last proteins down-regulated during the process of phenotypic modulation. Smoothelin is another smooth muscle contraction protein that is present in human and pig bladder detrusor

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muscle.41 Smoothelin is a cytoskeletal protein specifically expressed in differentiated SMCs and has been shown to co-localize with α-SM actin. Smooth muscle myosin heavy chains (SMHCs) are motor proteins that power smooth muscle contraction. SMHCs have two main isoforms: SM1 and SM2. SM1 is expressed in SMCs throughout embryonic development to the mature stage, whereas SM2 has a distinct function for smooth muscle contraction and is a specific marker of mature SMCs. The SM1/SM2 ratio is about 1 : 142 in normal bladder muscle tissue but decreases in overactive bladder.43 SM1 and SM2 isoforms of SMHC molecules are specific markers in immunostaining SMC differentiation. SMCs contain less protein than a typical striated muscle cell and much less myosin. The actin content is similar, so the ratio of actin to myosin is ∼6 : 1 in striated muscle and ∼15 : 1 in smooth muscle. Desmin reacts with parts of the intermediate filament muscle cell protein. In smooth muscle, desmin is located in the cytoskeletal region at the dense bodies and dense plaques. In the static culture environment, SMCs gradually lose the contractile phenotype as expression of genes such as desmin and myosin decreases in long-term culture.22,44–46 However, it has been shown that the presence of insulin-like growth factor, laminin,47 or transforming growth factor-β (TGF-β) and mechanical strain48 can prevent the spontaneous loss of the contractile phenotype of SMCs in vitro. Calponin acts as one regulatory contractile protein and has three isoforms: basic (H1, or α), neutral (H2), and acidic. H1-calponin is expressed only in differentiated SMCs.49 A lower molecular weight variant of calponin is often observed in SMCs of the human urogenital tract.50 Caldesmon, which has two isoforms, (H) and (L), is located in the thin filaments of SMCs and is a calmodulin- and actin-binding regulatory protein. H-caldesmon is predominantly expressed in differentiated SMCs, while L-caldesmon is expressed in dedifferentiated muscle cells. Both isoforms can be interconverted when cell phenotypes modulate.51 Vimentin is expressed in SMCs of the fetal bladder and its expression declines markedly during the postnatal period and during physiological hypertrophy of SMCs, which occurs concomitantly with decreases in DNA synthesis. Vimentin also exists in cytoplasm of myo-fibroblasts. There is an organized cytoskeleton consisting of the intermediate filament proteins vimentin and desmin, along with actin filaments.

16.2.2 Cells of non-urological origin In some cases, normal urological tissues are in short supply or unavailable for use in tissue engineering applications. In these cases, non-urological cells and tissues have been considered as an alternative for urological tissue regeneration. Several types of non-urological cells have been under investigation for urological tissue engineering, such as epithelium from buccal

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mucosa52–62 for urethra or bladder reconstruction and chrondocytes 63–68 and blood cells69 for correction of VUR. Epithelial cells and fibrolasts from skin have also been used for bladder regeneration70,71 (see Table 16.1).

16.2.3 Uses for differentiated cells in reconstruction of the urinary tract Bladder reconstruction has been well-investigated with tissue engineering technology. In 1999, the Atala group reported the results of a comparison study in which biodegradable scaffolds were either seeded with cells or used without cells for extensive bladder replacement procedures (>75%) in a canine model of subtotal cystectomy.4 A third animal group receiving cystoplasty surgery alone served as the control group. Animals underwent a subtotal cystectomy and were repaired with either a bladder-shaped polylactic-co-glycolic acid (PLGA)-based biodegradable scaffold that was seeded with autologous urothelium and SMCs or with a non-seeded bladder-shaped biodegradable scaffold. The cell-seeded scaffold performed better, creating a larger bladder capacity and nearly normal bladder compliance as measured by urodynamic studies. Additionally, histology data showed smooth muscle bundle formation and normal cellular organization with urothelial, suburothelial, and muscle layers within the regenerated bladder 9 months after bladder augmentation. In contrast, augmentation with unseeded scaffolds led to poor bladder tissue development, fibrosis, graft contraction, small bladder volume, and bladder dysfunction. Bladder volume remained small in the control group as well. This study demonstrated that cell-seeded scaffolds can improve bladder regeneration in models of extensive tissue replacement, but non-seeded scaffolds produce inadequate results in this case. Recently, another independent study demonstrated that autologous cells are necessary to achieve improved bladder tissue function when an extensive amount of bladder tissue is required.72 Further validation of the cell-seeded scaffold technology has been demonstrated in a clinical trial involving bladder augmentation in seven patients needing cystoplasty. Using autologous bladder biopsies, urothelial and muscle cells were grown in culture and seeded on to biodegradable bladdershaped scaffolds consisting of either collagen or a composite of collagen and polyglycolic acid (PGA). The seeded scaffolds were maintained in culture for 7 weeks. Then, autologous engineered bladder constructs were used for bladder reconstruction and implanted either with or without an omental wrap. At 2–7 years after surgery, the functional outcomes were greatest in the composite scaffold group, with no adverse events reported.7 Although cell-seeded scaffolds have an advantage over non-seeded technology in these cases, the mechanism governing this is not yet clear. It is uncertain whether the seeded cells proliferate and populate the scaffold themselves, or if they act in a ‘feeder cell-like’ manner and recruit and

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activate the local progenitor cells, which then complete the tissue regeneration process. Alternatively, the seeded cells could perform both of these functions. However, it is known that, for hollow organ tissue engineering, the bio-scaffold selected is very important. The ideal bio-scaffolds for regeneration of bladder or other hollow organs should possess the proper biomechanical and physical properties to allow the organ to develop and form the correct shape without collapsing before tissue regeneration is complete. For replacement of large amounts of tissue, PGA–PLGA is preferred. It can provide a three-dimensional space to allow the loading of more cells during the in vitro seeding phase, and this can lead to better tissue formation in vivo. Importantly, PGA–PLGA provides adequate physical properties to hold the shape of the implant and prevent collapse during tissue healing processes. Additionally, this material is highly porous, and this facilitates gas and nutrition exchange and promotes cell metabolism within the cell–scaffold composite. Finally, PGA–PLGA is already Food and Drug Administration (FDA) approved as a scaffold material, and has been used for many applications

16.3

Stem/progenitor cells for urological reconstruction

16.3.1 Stem/progenitor cells derived from the urinary tract Adult stem cells are defined as clonogenic, self-renewing progenitor cells that can generate one or more specialized types of cells. The majority of adult tissues have their own stem/progenitor cells that are capable of maintaining, generating, and replacing terminally differentiated cells within their own specific tissue in response to physiological cell turnover or tissue damage due to injury. Adult human bladder stem/progenitor cells have been recently described and characterized.12,25,73 In a recent study,25 we demonstrated preliminary evidence that single cells obtained from urine and bladder washes can be expanded to generate a diverse population of cells that express urothelial, muscle, endothelial, and interstitial cell markers. Three types of cells exist in urine: differentiated, differentiating, and progenitor cells. Most cells in urine are fully differentiated. They do not attach to tissue culture plates. About 0.1% of cells in urine are differentiating cells, which do attach to plates and display the morphology and protein markers of various bladder cell lineages. However, these cells do not expand further after subculture. About 0.2% of the cells in urine have a phenotype consistent with progenitor cells, which we termed ‘urine-derived progenitor cells’ (UPCs). These UPCs are easily cultured, appear genetically stable after a number of passages, and maintain the ability to give rise to more differentiated progeny cells. They express stem/progenitor cell markers

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such as c-Kit, SSEA4, and CD 44, and possess features of multipotent progenitor cells. In particular, they are capable of extensive expansion, a defining property of progenitor cells. UPCs give rise to multiple lineages that express urothelial, smooth muscle, endothelial, and interstitial cell markers in vitro. However, the extent of the self-renewal capability of UPCs remains to be established, in order to determine whether some members of the population may be classified as true stem cells. The number of UPC clones is on average about 2.5 cells/100 ml urine (from 2 to 7 cells/100 ml urine). We have shown that, a few days after being placed in a tissue culture well, a single cell forms a cluster of cells which appears small, compact and uniform. A consistently high yield of cells was achieved from each of these clonal lines. The cells reached confluence in about 2 weeks when placed in a 3-cm-diameter well at passage one. At passage two, cells were plated in 10 cm culture dishes and a cell number of approximately 1 million was reached in 3.5 weeks. Finally, in 6–7 weeks, the cultures expanded to approximately 100 million cells at passage four. These cells displayed normal exponential cell growth patterns, with a steady increase in cell numbers during a 10-day culture period. The average population doubling time was 31.3 hours in mixed media. These urine-derived cells also showed the ability to differentiate into various cell lineages as described below, and were capable of growing at least eight passages in vitro. We therefore consider them good candidates for progenitor cells and possibly stem cells. UPC clones were obtained from 40 to 70% of urine samples. Fresh urine gave the highest rate of colony formation (67%) and urine stored at 4 ºC the lowest (30%). Urine from volunteers aged 13–40 years gave the highest rate of clone recovery. Catheterization significantly enhanced the number of UPCs in urine compared with spontaneously voided urine, possibly because catheterization resulted in cells being scraped off the inner bladder wall. Collecting triple urine samples increased the rate of clone formation. There are many potential advantages to using UPCs as a cell source for urological tissue engineering. First, cells can be easily harvested and grown in culture. UPCs do not require enzyme digestion or culture on a layer of feeder cells to support cell growth. Second, since invasive surgical biopsy procedures are not necessary to harvest cells from urine, patient morbidity and potential complications, such as urethral or bladder trauma and urinary tract infections, are avoided. As UPCs are autologous somatic cells, no ethical issues are involved in their use for tissue reconstruction, and no immune reaction to engineered implants should occur. The quality of cells obtained from urine is similar to that of biopsyderived cells. When differentiated, UPCs express all of the proteins characteristic of the various bladder cell lineages. Karyotype analysis has demonstrated that these cells are genetically stable. Importantly, there is a

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major cost advantage to using UPCs – it costs about $50 to obtain cells from urine, compared with about $5000 to isolate cells from a biopsy procedure. About 1.4 × 109 urothelium cells and SMCs are required for bladder tissue regeneration.1 We estimate that three to four urine samples (about 15 UPCs/600 ml urine) expanded for 6–7 weeks would yield a sufficient quantity of low-passage, healthy cells for clinical tissue engineering applications. This time frame is comparable with that required for expansion from a tissue biopsy (7–8 weeks).1 UPCs and the cells obtained through urological tissue biopsies come from the same urinary tract systems and have similar biological features. Therefore, collecting cells from urine could be an attractive alternative to the standard urological tissue biopsies currently used in cell therapy and tissue engineering.

16.3.2 Stem/progenitor cells derived from non-urological tissues Despite the convenience of using differentiated cells in tissue engineering applications, these cells have several shortcomings. These cells have a limited ability to grow in culture and they tend to dedifferentiate in vitro, which may lead to insufficient numbers of cells. In addition, autologous bladder cells cannot be taken from patients with urinary tract malignancies. One solution to these problems is to prepare engineered tissues using stem cells. Bone marrow stem cells Currently, the most effectively characterized types of multipotent stem cells are from bone marrow. BMSCs have been shown to differentiate into specialized cells, including pneumocysts,9 hepatocytes,10 neural cells,11 skin,15 and ‘mesodermal derivatives’ such as bone, cartilage, cardiac muscles, skeletal muscle,12 and fat. When implanted BMSCs are placed on a proper bio-degradable scaffold, they can act as anti-fibrotic, angiogenic, antiapoptotic, and mitotic agents.74 Recently, BMSCs were evaluated as an alternative to replace bladder SMCs when native bladder muscle tissue is unavailable. The potential of BMSCs to differentiate into cells with bladder SMC characteristics was assessed in vitro and in different animal models.5,6,75 Kanematsu et al.75 showed that in vitro, when either supernatants from cultured bladder cells (conditioned media) or media containing TGF-β and vascular endothelial growth factor (VEGF) induced bone marrow cells to adopt an SMC phenotype. In order to test this in vivo, bone marrow cells expressing green fluorescent protein were transplanted into lethally irradiated rats. Eight weeks following transplantation, bladder domes were replaced with bladder

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acellular matrix grafts. Two weeks after the graft procedure, green fluorescent protein (GFP) expression in the matrices indicated that the transplanted marrow cells had repopulated the graft. By 12 weeks, these cells reconstituted the smooth muscle layer, with native SMCs also infiltrating the graft. In another rat study,5 rapid bladder SMC and urothelium regeneration occurred on BMSC-seeded collagen matrices, whereas fibrotic changes were observed in the non-seeded matrix group 3 months after bladder augmentation. In a large animal study,6 BMSCs proliferated at the same rate as primary cultured bladder SMCs in vitro, and they had a similar histological appearance and contractile phenotype to primary cultured bladder SMCs. BMSCs had a significant contractile response to calcium ionophore in vitro which was similar to that of bladder SMCs but markedly different to that of fibroblasts. Immunohistochemical staining and Western blotting indicated that BMSCs expressed α-SM actin, but did not express desmin or myosin. In vivo, small intestinal submucosa (SIS) grafts seeded with BMSCs developed solid smooth muscle bundle formations throughout the grafts, as did bladder cell-seeded SIS grafts. However, bladder tissue regeneration did not occur in the animal group that received cell-free scaffolding. The results indicate that BMSCs may provide an alternative cell source in bladder tissue engineering, which is important for bladder cancer patients who require bladder augmentation or replacement, but do not have enough normal, non-malignant bladder cells to use in tissue engineering applications. Other types of stem/progenitor cells have been studied as cell sources for bladder regeneration and cell therapy for stress urinary incontinence. These include skeletal muscle-derived progenitor cells and adipose stem cells.60

16.3.3 Methods of inducing differentiation of stem/ progenitor cells It is desirable to mimic the physiological conditions that guide stem cells to differentiate into the desired target cells before implantation. Several factors have been shown to enhance autologous adult stem cell differentiation into functional SMCs, and these are outlined below. 1

Growth factors, such as VEGF, platelet-derived growth factor (PDGFbb), TGF-β, and insulin-like growth factor (IGF).47 2 Components of the extracellular matrix (ECM). Cellular interactions with the ECM play an important role in cell adhesion, growth, migration, apoptosis, and differentiation.76 The ECM consists of compounds such as collagen, laminin,77 and fibronectin. Collagen IV can promote embryonic stem cells to differentiate into stem cell antigen-1-positive (Sca-1+) progenitor cells and SMCs.78 Additionally, culture on a three-

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dimensional ECM scaffold in a dynamic culture system can improve cell proliferation and differentiation and lead to a more homogeneous distribution of cells on the scaffold79,80 when compared with twodimensional static culture. 3 In vitro co-culture of fully differentiated SMCs and stem cells appears to improve stem cell differentiation into muscle cells, most likely because the SMCs secrete specific factors into the culture medium. Baskin’s group81 demonstrated that mature urothelium can induce urological embryonic tissue or stem cells to differentiate into SMCs in vivo through epithelial–stromal cell interaction or cell–cell interaction. In contrast, the embryonic tissue failed to differentiate into SMCs when urothelium was not present.81 Like co-culture, conditioned medium is commonly used for stem cell differentiation. Conditioned medium is a supernatant of culture medium that is partially used by cells and, because of this, it is enriched with cell-derived material including small amounts of growth factors. 4 The application of cyclic mechanical strain to a culture has been demonstrated to increase the expression of SMC markers in stem cells.82,83 Periodic stretching occurs in vivo as a part of the natural function of hollow organs; for example, as the bladder fills and empties. Differentiated SMCs easily lose their contractile function in static culture once the cells leave the body, the use of mechanical strain in culture can prevent the spontaneous loss of the contractile phenotype of SMCs in vitro and maintain SMC functional characteristics.84

16.3.4 Autologous stem cells for endoscopic therapies Another exciting area of clinical urological investigation is the use of various autologous cells in cell therapy for VUR and USI. An increasing number of clinical trials using tissue engineering approaches have been reported (Table 16.1). All clinical applications of urological tissue engineering are based on a series of successful animal experiments.4,27,67,68 Endoscopic therapy offers a simple method for definitive treatment in USI and VUR. Two types of injectable substances have been investigated. First, natural and synthetic biomaterials that serve as bulking agents – such as silicon, fibrin, bioglass, polyvinyl alcohol foam, alginate gel, a smallintestinal submucosal suspension, and Deflux – have been used.85 Currently, injectable therapy based on bulking agents is used for only about two-thirds of patients with USI, with even lower cure rates. Potential problems of these injectable substances include the need for multiple injections to obtain and maintain optimal efficacy, potential antigenicity of the injectable and related allergic reaction, migration of injected material, and urethral pain both at the time of injection and afterward. The ideal bulking agent should remain

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efficacious over time and have few side effects. So far, none of these substances has met these criteria for success and the search for a superior injectable therapy for USI continues. Cell-based therapy is a promising alternative in urological procedures for VUR and USI. Autologous cells that can be used for this purpose include chondrocytes,67–69 adipose-derived stem cells (ASCs),26–28 BMSCs,6,29 and skeletal muscle-derived progenitor cells.31,32 In 1994, Atala et al.67,68 proposed the use of injectable autologous chondrocytes to correct VUR via endoscopy. In a minipig model, they noted that chondrocytes injected around the ureters did not migrate and the cartilage bead produced by this technique maintained its volume with time. Reflux was corrected in all animals treated endoscopically with autologous chondrocytes. After successful animal experiments, several clinic trials have been started.63,65 The new study63 was conducted in a total of 29 children (46 ureters) with grades II to IV reflux. Chondrocytes were harvested from a biopsy of each patient’s ear cartilage and were grown in culture for 6 weeks. Patients then returned to the clinic for transurethral injection of chondrocytes into the bladder trigone to correct reflux. Ultrasound was performed 1 month postoperatively and radionuclide cystography was carried out 3 months postoperatively to confirm reflux resolution. When reflux persisted, repeat treatment with stored chondrocytes was offered. In this study, a single chondrocyte injection corrected reflux in 26 of the 46 ureters (57%), while secondary injection was successful in 12 of 19 (63%). Overall, reflux was corrected in 38 of the 46 ureters (83%) and in 24 of the 29 patients (83%). There were no significant complications, and transurethral injection of autologous chondrocytes to correct VUR in children appears to be an effective and safe technique. The only limitation of this therapy is the high cost. Myocyte- or stem cell-based injection therapies have also been tested in VUR cases. Primary VUR is a congenital anomaly of the ureter–vesical junction. This creates a deficiency of the longitudinal muscle of the intravesical ureter, which leads to an inadequate valvular mechanism and allows urine to flow backward from the bladder to the kidney. Thus, muscle-based therapies are attractive options for the recovery of this muscle defect at the ureteral orifice. Autologous progenitor cell-based therapy has also made significant progress in treatment of USI. This cell therapy could soon become a standard procedure. The objective of this therapy is to improve or cure the sphincter dysfunction via periurethral endoscopic injection. Currently, myoblasts obtained from skeletal muscle biopsies86 and adipose-derived cells are the most commonly used cells for therapy for USI. Recently, autologous myoblasts and fibroblasts have been evaluated as a potential injectable therapy for USI. One group 86 has studied a combination therapy consisting of autologous myoblasts injected into the rhabdosphincter and fibroblasts

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injected into the urethral submucosa. A long-term study of 121 women and 63 men was performed from 2002 to 2005. With a minimum follow-up of 1 year, a cure rate of about 80% with improvements in quality-of-life scores, rhabdosphincter contractility, and urethral closure pressures has been achieved. All patients were continent 1 year after receiving this therapy and maintained their good outcome at further follow-up visits 3 and 4 years post-treatment. Ultrasound images before treatment clearly revealed poor periurethral integrity of the sphincteric mechanism; postinjection images revealed a completely normal-appearing urethra. Rodriguez et al.28 recently reported that adipose-derived stem cells (ASCs) have the potential to differentiate into functional SMCs. ASCs expressed a series of contractile proteins – including α-SM actin, desmin, myosin heavy chain, calponin, caldesomon, smoothlin, and SM22 – following aspiration from fat tissue and culture in smooth muscle differentiation medium. One important advantage of using ASCs is that adipose tissue can be harvested in large quantities with minimal morbidity. Autologous fat tissues were used for cell injection therapy for vesicoureteral reflux in a clinical study. Out of 11 patients, 2 had a reduction in grade of reflux, including one ureter that ceased refluxing altogether.26 One recent study87 showed that ASCs could correct neurogenic erectile dysfunction in rats as effectively as BMSCs. More research is underway to determine whether ASCs can differentiate into Leydig, Sertoli, and male germ cells. The eventual goal of the research is to use ASCs to treat male infertility and testosterone deficiency.

16.4

Cell tracking technology

Cell tracking technology has become necessary in order to monitor the fate (viability, differentiation, clonogenicity, and function) of grafted cells following implantation. These cells must be identified and localized within the host environment. Tracking provides information about how the graft cells integrate into the new setting and may even provide information about their vital functions. Several tracking approaches have been commonly used to identify graft cells: (a) prelabeling techniques, which rely on markers that were introduced into the cells prior to transfer into recipient animals; (b) postlabeling techniques, which depend on the detection of cell type-specific proteins that are analyzed in situ or after re-isolation of the cells; (c) the presence of genetic markers (Table 16.2). An ideal tracking marker should label cells efficiently and stably. It should not interfere with any cellular functions and it should not provoke immune reactions. In addition, it should ideally allow the specific and sensitive detection of single cells in vivo, as well as the re-isolation of labeled cells for functional analysis. However, most current markers only partially

Genetic marker

Post-labeling

Pre-labeling

Tracking approaches

Whole cell

Molecular imaging

Y-chromosome

Y-chromosome (only male)

Mitochrome

Whole cell

LacZ

Human mitochromes

Cytoplasm

GFP

Nuclear

DNA

Bromodesoxyuidine (BrdU)

Human nuclear staining

Membrane

PKH 2, 3 and 26 fluorescent dye

Markers

Sub-cellular localization

Tracking donorderived cells in vivo

Monitor the viability and gene expression of transplanted cells in vivo Monitor the viability and gene expression of transplanted cells in vivo Magnetically labeled cells follow and track cells in vivo by magnetic resonance imaging (MRI) Identify human cells in xenograft models Tracking human grafted cells

Estimation of cell life span; study cellular proliferation in vivo

Determine cell migration and proliferation in vivo

Purposes

Table 16.2 Methods of tracking grafted cells and cellular markers

Stable labeling; not tissue-specific; indentify donor male nuclei

Reacts specifically with human cells

Reacts specifically with human cells

Non-invasive monitoring of cellular movements

Stable and reliable long-term labeling; do not affect cellular function; track stem cell differentiation Stable and reliable long-term labeling

PKH 26: not cytotoxic; stable labeling for up to 3 months; intense and easy to use Loss of the labeling intensity as the cell divides

Advantages

Fetidex loading might block cellular differentiation of stem cells; weak signal from the labeling cells in deeper tissue in larger animal models Human mitochrome expression screens are particularly laborious Human mitochrome expression screens are particularly laborious Implantation of male graft tissue to female animal might cause immune response against male cells

LacZ is heat labile, and the enzyme activity is destroyed during paraffin embedding

PKH 2: elutes and interferes with proliferation and migration; minimal leaking from cell to cell Toxic for the replication of some cells; not for the long-term follow-up of labeling cells Free GFP is extremely soluble and leaks out from liquid-covered cryostat sections

Limitations

Properties

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fulfill these criteria. Identification of transplanted cells has been generally difficult. Choosing the right marker depends on the experimental set-up, the type of tissue manipulation needed for marker detection, and the stability of the marker in vivo. Many markers are commercially available – including specific DNA sequences, proteins, dyes, or radioactive metabolites – and these can be targeted to the membrane, the cytoplasm, or the nucleus of the target cells. Visualization of these markers usually requires specific detection techniques such as antibody staining, in situ hybridization, or autoradiography, unless fluorochromes are used that are detectable without further manipulation of the labeled cells. Three biomarkers are commonly used in tracking grafted cells: GFP, LacZ, and a Y-chromosome (male)-specific probe. GFP is a protein isolated from deep sea jellyfish (Aequorea victoria). Many variations of the protein are now available, providing a wide range of fluorescent colors to use as different markers. These GFP reporter genes have now been introduced into urothelial cells along with the endothelial growth factor gene,88 which was intended to improve blood supply for urethral repair. A major limitation of using GFP to measure gene expression within regenerating tissue is that the grafted tissues may exhibit high intensity autofluorescence, leading to high background readings during measurement and difficulty in interpreting the data. Sensitivity of detection is also an issue for quantitative assays since GFP is not as sensitive when used to measure gene expression at the chromosomal level. Additionally, the instability and degradation of GFP can be an issue when looking at gene expression events in real time. Reporter genes such as LacZ, which encodes the bacterial enzyme β-galactosidase, can also be used. Cells expressing the enzyme can be visualized using a variety of techniques. Most involve cleavage of specific substrates by β-galactosidase; these substrates consist of galactose linked to a moiety whose properties change upon enzymatic liberation from galactose. Several substrates yield colored or fluorescent soluble products, which are useful when quantifying β-galactosidase activity or visualizing transduced cells live in vivo. For localization of cells containing transduced LacZ in situ in tissue sections, chromogenic substrates that yield a precipitated product are desirable. Yakoyama et al.89,90 reported that LacZ-labeled skeletal muscle-derived progenitor cells persist and survive in a rat model of USI treatment 1 month after injection. In addition, Y-chromosome (male)-specific probes can be used to identify and track transplanted cells in a sex-mismatched cell transplantation model. In mammals, the Y chromosome is present only in the nucleus of a healthy male cell. Thus, the Y chromosome is an excellent marker for detection of donor male nuclei after implantation into female animals.91 It is not tissuespecific, does not require transfection or measurement of gene expression, and the staining is stable. However, localization of transplanted cells with

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X and Y chromosome prober is limited by the sparse signals generated. Quantifying total male DNA in samples is feasible using slot blots or quantitative reverse transcriptase-PCR (QRT-PCR or qRT-PCR). Other markers, such as anti-human nuclei staining,92,93 human nuclear matrix91 and human microchromal antigen,92 pan-nuclear signals with a total human DNA probe, or multiple nuclear signals with a pan-centromeric human DNA probe,94 have been used in cell tracking studies. In addition, recent advances in the field of molecular imaging have provided noninvasive means to track the effects of cell-based therapy.95 Non-invasive techniques have been explored for in vivo cell monitoring after cellular implantation (see Table 16.2). This tracking technology can bridge the gap between the extensive knowledge of cell-cell interaction in vitro and the imaging techniques that can reveal dynamic processes in vivo in stem cell biology, immunology, and development.

16.5

Conclusions

Current advances in urological tissue engineering and cell-based therapy demonstrate that bladder and urethral tissues can be regenerated using autologous cells seeded onto biodegradable scaffolds. VUR and USI can be corrected with injections of autologous cells contained in a hydrogel. However, many issues must be elucidated before these techniques can become widely used in the clinic. For example, the role of donor cells in tissue regeneration remains unclear, and it is not known whether the seeded cells proliferate and populate scaffold materials themselves, or if they stimulate the activation, migration, proliferation, and differentiation of the local progenitor cells to complete the tissue regeneration. In addition, an approach to promote angiogenesis and to facilitate innervation with a functional network of regenerated nerves will greatly improve tissue regeneration strategies to create a de novo urological organ.

16.6

Acknowledgements

The author would like to thank Ms Karen Klein and Dr Jennifer Olson for editorial assistance.

16.7

References

1 atala, a.: Advances in tissue and organ replacement. Curr Stem Cell Res Ther, 3: 21, 2008. 2 atala, a.: Tissue engineering, stem cells, and cloning for the regeneration of urologic organs. Clin Plast Surg, 30: 649, 2003. 3 yoo, j. j., meng, j., oberpenning, f. et al.: Bladder augmentation using allogenic bladder submucosa seeded with cells. Urology, 51: 221, 1998.

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4 oberpenning, f., meng, j., yoo, j. j. et al.: De novo reconstitution of a functional mammalian urinary bladder by tissue engineering. Nat Biotechnol, 17: 149, 1999. 5 chung, s. y., krivorov, n. p., rausei, v. et al.: Bladder reconstitution with bone marrow derived stem cells seeded on small intestinal submucosa improves morphological and molecular composition. J Urol, 174: 353, 2005. 6 zhang, y., lin, h. k., frimberger, d. et al.: Growth of bone marrow stromal cells on small intestinal submucosa: an alternative cell source for tissue engineered bladder. BJU Int, 96: 1120, 2005. 7 atala, a., bauer, s. b., soker, s. et al.: Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet, 367: 1241, 2006. 8 zhang, y., frimberger, d., cheng, e. y. et al.: Challenges in a larger bladder replacement with cell-seeded and unseeded small intestinal submucosa grafts in a subtotal cystectomy model. BJU Int, 98: 1100, 2006. 9 kropp, b. p., ludlow, j. k., spicer, d. et al.: Rabbit urethral regeneration using small intestinal submucosa onlay grafts. Urology, 52: 138, 1998. 10 badylak, s. f., kropp, b., mcpherson, t. et al.: Small intestional submucosa: a rapidly resorbed bioscaffold for augmentation cystoplasty in a dog model. Tissue Eng, 4: 379, 1998. 11 kropp, b. p., rippy, m. k., badylak, s. f. et al.: Regenerative urinary bladder augmentation using small intestinal submucosa: urodynamic and histopathologic assessment in long-term canine bladder augmentations. J Urol, 155: 2098, 1996. 12 oottamasathien, s., wang, y., williams, k. et al.: Directed differentiation of embryonic stem cells into bladder tissue. Dev Biol, 304: 556, 2007. 13 frimberger, d., morales, n., shamblott, m. et al.: Human embryoid bodyderived stem cells in bladder regeneration using rodent model. Urology, 65: 827, 2005. 14 lakshmanan, y., frimberger, d., gearhart, j. d. et al.: Human embryoid bodyderived stem cells in co-culture with bladder smooth muscle and urothelium. Urology, 65: 821, 2005. 15 atala, a.: Autologous cell transplantation for urologic reconstruction. J Urol, 159: 2, 1998. 16 atala, a.: Future perspectives in reconstructive surgery using tissue engineering. Urol Clin North Am, 26: 157, 1999. 17 atala, a.: New methods of bladder augmentation. BJU Int, 85 Suppl 3: 24, 2000. 18 atala, a.: Future trends in bladder reconstructive surgery. Semin Pediatr Surg, 11: 134, 2002. 19 atala, a.: Future perspectives in bladder reconstruction. Adv Exp Med Biol, 539: 921, 2003. 20 atala, a.: Recent developments in tissue engineering and regenerative medicine. Curr Opin Pediatr, 18: 167, 2006. 21 atalan, g., cihan, m., sozmen, m. et al.: Repair of urethral defects using fascia lata autografts in dogs. Vet Surg, 34: 514, 2005. 22 kropp, b. p., zhang, y., tomasek, j. j. et al.: Characterization of cultured bladder smooth muscle cells: assessment of in vitro contractility. J Urol, 162: 1779, 1999. 23 zhang, y., kropp, b. p., moore, p. et al.: Coculture of bladder urothelial and smooth muscle cells on small intestinal submucosa: potential applications for tissue engineering technology. J Urol, 164: 928, 2000.

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24 zhang, y. y., bailey, r. r.: Treatment of vesicoureteric reflux in a sheep model using subureteric injection of cultured fetal-bladder tissue. Pediatr Surg Int, 13: 32, 1998. 25 zhang, y., mcneill, e., tian, h., et al.: Urine-derived cells are a potential source for urologic tissue reconstruction. J Urol, 180: 2226, 2008. 26 palma, p. c., ferreira, u., ikari, o. et al.: Subureteric lipoinjection for vesicoureteral reflux in renal transplant candidates. Urology, 43: 174, 1994. 27 palma, p. c., vidal, b., riccetto, c. l. et al.: Effect of purified collagen on lipograft survival: experimental basis for periurethral lipoinjections. J Endourol, 17: 255, 2003. 28 rodriguez, l. v., alfonso, z., zhang, r. et al.: Clonogenic multipotent stem cells in human adipose tissue differentiate into functional smooth muscle cells. Proc Natl Acad Sci USA, 103: 12167, 2006. 29 de coppi, p., callegari, a., chiavegato, a. et al.: Amniotic fluid and bone marrow derived mesenchymal stem cells can be converted to smooth muscle cells in the cryo-injured rat bladder and prevent compensatory hypertrophy of surviving smooth muscle cells. J Urol, 177: 369, 2007. 30 kameda, t., hatakeyama, s., terada, k. et al.: Acceleration of the formation of cultured epithelium using the sonic hedgehog expressing feeder cells. Tissue Eng, 7: 545, 2001. 31 li, y., huard, j.: Differentiation of muscle-derived cells into myofibroblasts in injured skeletal muscle. Am J Pathol, 161: 895, 2002. 32 lu, s. h., cannon, t. w., chermanski, c. et al.: Muscle-derived stem cells seeded into acellular scaffolds develop calcium-dependent contractile activity that is modulated by nicotinic receptors. Urology, 61: 1285, 2003. 33 southgate, j., hutton, k. a., thomas, d. f. et al.: Normal human urothelial cells in vitro: proliferation and induction of stratification. Lab Invest, 71: 583, 1994. 34 zhang, y. y., frey, p.: Growth of cultured human urothelial cells into stratified urothelial sheet suitable for autografts. Adv Exp Med Biol, 539: 907, 2003. 35 sugasi, s., lesbros, y., bisson, i. et al.: In vitro engineering of human stratified urothelium: analysis of its morphology and function. J Urol, 164: 951, 2000. 36 lobban, e. d., smith, b. a., hall, g. d. et al.: Uroplakin gene expression by normal and neoplastic human urothelium. Am J Pathol, 153: 1957, 1998. 37 zhang, y., kropp, b. p., lin, h. k. et al.: Bladder regeneration with cell-seeded small intestinal submucosa. Tissue Eng, 10: 181, 2004. 38 de filippo, r. e., yoo, j. j., atala, a.: Urethral replacement using cell-seeded tubularized collagen matrices. J Urol, 168: 1789, 2002. 39 chamley-campbell, j., campbell, g. r., ross, r.: The smooth muscle cell in culture. Physiol Rev, 59: 1, 1979. 40 koh, g. y., klug, m. g., soonpaa, m. h. et al.: Differentiation and long-term survival of C2C12 myoblast grafts in heart. J Clin Invest, 92: 1548, 1993. 41 maake, c., landman, m., wang, x. et al.: Expression of smoothelin in the normal and the overactive human bladder. J Urol, 175: 1152, 2006. 42 dotsch, j., hanze, j., knufer, v. et al.: Increased urinary adrenomedullin excretion in children with urinary-tract infection. Nephrol Dial Transplant, 13: 1686, 1998. 43 fitzgerald, m. p., manaves, v., martin, a. f. et al.: Myosin isoforms in female human detrusor. Neurourol Urodyn, 20: 23, 2001.

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44 chamley, j. h., campbell, g. r.: Mitosis of contractile smooth muscle cells in tissue culture. Exp Cell Res, 84: 105, 1974. 45 groschel-stewart, u., chamley, j. h., campbell, g. r. et al.: Changes in myosin distribution in dedifferentiating and redifferentiating smooth muscle cells in tissue culture. Cell Tissue Res, 165: 13, 1975. 46 panettieri, r. a., murray, r. k., depalo, l. r. et al.: A human airway smooth muscle cell line that retains physiological responsiveness. Am J Physiol, 256: C329, 1989. 47 hayashi, k., saga, h., chimori, y. et al.: Differentiated phenotype of smooth muscle cells depends on signaling pathways through insulin-like growth factors and phosphatidylinositol 3-kinase. J Biol Chem, 273: 28860, 1998. 48 stegemann, j. p., nerem, r. m.: Phenotype modulation in vascular tissue engineering using biochemical and mechanical stimulation. Ann Biomed Eng, 31: 391, 2003. 49 walsh, m. p.: Calponin–knocked out but not down! J Physiol, 529 Pt 3: 517, 2000. 50 draeger, a., gimona, m., stuckert, a. et al.: Calponin. Developmental isoforms and a low molecular weight variant. FEBS Lett, 291: 24, 1991. 51 sobue, k., hayashi, k., nishida, w.: Expressional regulation of smooth muscle cell-specific genes in association with phenotypic modulation. Mol Cell Biochem, 190: 105, 1999. 52 mokhless, i. a., kader, m. a., fahmy, n. et al.: The multistage use of buccal mucosa grafts for complex hypospadias: histological changes. J Urol, 177: 1496, 2007. 53 mehrsai, a., djaladat, h., salem, s. et al.: Outcome of buccal mucosal graft urethroplasty for long and repeated stricture repair. Urology, 69: 17, 2007. 54 mclaughlin, m. d., thrasher, j. b., celmer, a. et al.: Buccal mucosal urethroplasty in patients who had multiple previous procedures. Urology, 68: 1156, 2006. 55 levine, l. a., strom, k. h., lux, m. m.: Buccal mucosa graft urethroplasty for anterior urethral stricture repair: evaluation of the impact of stricture location and lichen sclerosus on surgical outcome. J Urol, 178: 2011, 2007. 56 el-kasaby, a. w., fath-alla, m., noweir, a. m. et al.: The use of buccal mucosa patch graft in the management of anterior urethral strictures. J Urol, 149: 276, 1993. 57 dessanti, a., rigamonti, w., merulla, v. et al.: Autologous buccal mucosa graft for hypospadias repair: an initial report. J Urol, 147: 1081, 1992. 58 burger, r. a., muller, s. c., el-damanhoury, h. et al.: The buccal mucosal graft for urethral reconstruction: a preliminary report. J Urol, 147: 662, 1992. 59 brock, j. w., 3rd: Autologous buccal mucosal graft for urethral reconstruction. Urology, 44: 753, 1994. 60 bhargava, s., chapple, c. r., bullock, a. j. et al.: Tissue-engineered buccal mucosa for substitution urethroplasty. BJU Int, 93: 807, 2004. 61 berglund, r. k., angermeier, k. w.: Combined buccal mucosa graft and genital skin flap for reconstruction of extensive anterior urethral strictures. Urology, 68: 707, 2006. 62 ahmed, s., gough, d. c.: Buccal mucosal graft for secondary hypospadias repair and urethral replacement. Br J Urol, 80: 328, 1997. 63 paltiel, h. j., diamond, d. a., zurakowski, d. et al.: Endoscopic treatment of vesicoureteral reflux with autologous chondrocytes: postoperative sonographic features. Radiology, 232: 390, 2004.

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64 lee, j. w., kim, y. h., kim, s. h. et al.: Chondrogenic differentiation of mesenchymal stem cells and its clinical applications. Yonsei Med J, 45 Suppl: 41, 2004. 65 diamond, d. a., caldamone, a. a.: Endoscopic correction of vesicoureteral reflux in children using autologous chondrocytes: preliminary results. J Urol, 162: 1185, 1999. 66 caldamone, a. a., diamond, d. a.: Long-term results of the endoscopic correction of vesicoureteral reflux in children using autologous chondrocytes. J Urol, 165: 2224, 2001. 67 atala, a., kim, w., paige, k. t. et al.: Endoscopic treatment of vesicoureteral reflux with a chondrocyte-alginate suspension. J Urol, 152: 641, 1994. 68 atala, a., cima, l. g., kim, w. et al.: Injectable alginate seeded with chondrocytes as a potential treatment for vesicoureteral reflux. J Urol, 150: 745, 1993. 69 sugiyama, t., hanai, t., hashimoto, k. et al.: Long-term outcome of the endoscopic correction of vesico-ureteric reflux: a comparison of injected substances. BJU Int, 94: 381, 2004. 70 brehmer, b., rohrmann, d., becker, c. et al.: Different types of scaffolds for reconstruction of the urinary tract by tissue engineering. Urol Int, 78: 23, 2007. 71 brehmer, b., rohrmann, d., rau, g. et al.: Bladder wall replacement by tissue engineering and autologous keratinocytes in minipigs. BJU Int, 97: 829, 2006. 72 jayo, m. j., jain, d., wagner, b.j.: Early cellular and stromal responses in regeneration versus repair of a mammalian bladder using autologous cell and biodegradable scaffold technologies. J Urol, 180: 392, 2008. 73 fierabracci, a., caione, p., di giovine, m. et al.: Identification and characterization of adult stem/progenitor cells in the human bladder (bladder spheroids): perspectives of application in pediatric surgery. Pediatr Surg Int, 23: 837, 2007. 74 caplan, a. i.: Adult mesenchymal stem cells for tissue engineering versus regenerative medicine. J Cell Physiol, 213: 341, 2007. 75 kanematsu, a., yamamoto, s., iwai-kanai, e. et al.: Induction of smooth muscle cell-like phenotype in marrow-derived cells among regenerating urinary bladder smooth muscle cells. Am J Pathol, 166: 565, 2005. 76 hirst, s. j., twort, c. h., lee, t. h.: Differential effects of extracellular matrix proteins on human airway smooth muscle cell proliferation and phenotype. Am J Respir Cell Mol Biol, 23: 335, 2000. 77 hubschmid, u., leong-morgenthaler, p. m., basset-dardare, a. et al.: In vitro growth of human urinary tract smooth muscle cells on laminin and collagen type I-coated membranes under static and dynamic conditions. Tissue Eng, 11: 161, 2005. 78 xiao, q., zeng, l., zhang, z. et al.: Stem cell-derived Sca-1+ progenitors differentiate into smooth muscle cells, which is mediated by collagen IV-integrin alpha1/beta1/alphav and PDGF receptor pathways. Am J Physiol Cell Physiol, 292: C342, 2007. 79 arrigoni, c., chitto, a., mantero, s. et al.: Rotating versus perfusion bioreactor for the culture of engineered vascular constructs based on hyaluronic acid. Biotechnol Bioeng, 100: 988, 2008. 80 gomes, m. e., sikavitsas, v. i., behravesh, e. et al.: Effect of flow perfusion on the osteogenic differentiation of bone marrow stromal cells cultured on starch-based three-dimensional scaffolds. J Biomed Mater Res A, 67: 87, 2003.

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81 disandro, m. j., li, y., baskin, l. s. et al.: Mesenchymal-epithelial interactions in bladder smooth muscle development: epithelial specificity. J Urol, 160: 1040, 1998. 82 kobayashi, n., yasu, t., ueba, h. et al.: Mechanical stress promotes the expression of smooth muscle-like properties in marrow stromal cells. Exp Hematol, 32: 1238, 2004. 83 huang, h., nakayama, y., qin, k. et al.: Differentiation from embryonic stem cells to vascular wall cells under in vitro pulsatile flow loading. J Artif Organs, 8: 110, 2005. 84 akhyari, p., fedak, p. w., weisel, r. d. et al.: Mechanical stretch regimen enhances the formation of bioengineered autologous cardiac muscle grafts. Circulation, 106: I137, 2002. 85 puri, p., chertin, b., velayudham, m. et al.: Treatment of vesicoureteral reflux by endoscopic injection of dextranomer/hyaluronic acid copolymer: preliminary results. J Urol, 170: 1541, 2003. 86 ginsberg, d.: Urinary incontinence: new data from AUA 2007, in Highlights of the American Urological Association (AUA) 2007 Annual Meeting, 2007, 19–24 May 2007, Anaheim, CA, AUA. 87 lin, c. s., xin, z. c., deng, c. h. et al.: Recent advances in andrology-related stem cell research. Asian J Androl, 10: 171, 2008. 88 guan, y., ou, l., hu, g. et al.: Tissue engineering of urethra using human vascular endothelial growth factor gene-modified bladder urothelial cells. Artif Organs, 32: 91, 2008. 89 yokoyama, t., yoshimura, n., dhir, r. et al.: Persistence and survival of autologous muscle derived cells versus bovine collagen as potential treatment of stress urinary incontinence. J Urol, 165: 271, 2001. 90 yokoyama, t., huard, j., pruchnic, r. et al.: Muscle-derived cell transplantation and differentiation into lower urinary tract smooth muscle. Urology, 57: 826, 2001. 91 beier, j. p., stern-straeter, j., foerster, v. t. et al.: Tissue engineering of injectable muscle: three-dimensional myoblast-fibrin injection in the syngeneic rat animal model. Plast Reconstr Surg, 118: 1113, 2006. 92 daadi, m. m., saporta, s., willing, a. e. et al.: In vitro induction and in vivo expression of bcl-2 in the hNT neurons. Brain Res Bull, 56: 147, 2001. 93 klein, s. m., behrstock, s., mchugh, j. et al.: GDNF delivery using human neural progenitor cells in a rat model of ALS. Hum Gene Ther, 16: 509, 2005. 94 cho, j. j., malhi, h., wang, r. et al.: Enzymatically labeled chromosomal probes for in situ identification of human cells in xenogeneic transplant models. Nat Med, 8: 1033, 2002. 95 kraitchman, d. l., tatsumi, m., gilson, w. d. et al.: Dynamic imaging of allogeneic mesenchymal stem cells trafficking to myocardial infarction. Circulation, 112: 1451, 2005. .

17 Embryonic stem cells, nuclear transfer and parthenogenesis-derived stem cells for urological reconstruction R. D O R I N, J. YA M Z O N and C. J. K O H , Childrens Hospital Los Angeles and University of Southern California Keck School of Medicine, USA

Abstract: Regenerative medicine has made substantial progress in addressing congenital and acquired diseases that result in end organ failure by combining the principles of cell biology, materials science, engineering, and transplantation. This chapter will discuss the basic principles of tissue engineering and the contemporary use of stem cells in the field of urological research. It will also provide a review of leading stem cell procurement techniques utilized in contemporary tissue engineering labs. These include embryonic stem cell harvest, parthenogenesis, nuclear transfer, and most recently, induced pluripotent stem cells. Key words: regenerative medicine, tissue engineering, embryonic stem cells, parthenogenesis, nuclear transfer, induced pluripotent stem cells.

17.1

Introduction

Regenerative medicine has made substantial progress in addressing congenital and acquired diseases that result in end organ failure. In combining principles of cell biology, materials science, engineering, and transplantation, many therapeutic strategies are under development to replace or rehabilitate diseased tissues. Traditional clinical applications of tissue replacement in urology include the use of autologous tissues, such as intestinal substitution in bladder augmentation, or the use of allogenic grafts, such as in kidney transplantation. Though these tried and tested approaches have successfully reduced the mortality and morbidity of the urological diseases to which they are applied, they are associated with many potential complications, making these types of tissue replacement less than ideal. When feasible, the use of autologous tissues is generally preferred to allogenic substitutes, which are complicated by the elicited host immune response. However, procurement of tissue from adjacent organs may incur complications and may not restore full natural function. For example, 357

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bladder augmentation for high-pressure neurogenic bladder is performed using autologous intestinal segments to achieve low-pressure storage with adequate capacity. However, interposition of intestinal segments into the urinary tract is associated with an increased risk of urolithiasis, metabolic disturbance, obstructive uropathy, and malignancy. Organ transplantation is complicated by the need for tissue matching, chronic immunosuppression, and the risk of organ rejection. Furthermore, the number of patients awaiting transplantation far surpasses the number of available donor organs. In the year 2004, approximately 472 100 patients were being treated for end stage renal disease in the United States, 336 000 of which were undergoing renal dialysis. Only 16 905 renal transplants were performed that year, and 84 252 patients died of their disease according to the National Kidney and Urologic Diseases Information Clearinghouse of the National Institutes of Health. By the year 2010, over 2 million adults and children are projected to suffer from end stage renal disease globally, with the projected associated cost of maintenance dialysis exceeding 1 trillion dollars (Lysaght and Hazlehurst, 2004). Advancements in medicine and surgery have increased the demand for the development of readily available, reproducible, non-immunogenic tissue sources for organ rehabilitation or replacement. Tissue engineering, a field within regenerative medicine, arose in the effort to meet these increasing demands. The most current and clinically applicable methods of tissue engineering utilize autologous tissues (Atala et al., 2006; De Filippo et al., 2002; De Filippo et al., 2003; Yoo et al., 1998). Cellular constructs created in this manner are patient specific and derived from a biopsy, necessitating an extra procedure. The ideal engineered tissue substitutes would be reproducible and applied universally without concern for immunogenicity. Tissue engineered from stem cells can potentially meet these criteria, and thus offer source material for organ reconstruction and regeneration without the need for additional biopsy procedures. This chapter will discuss the basic principles of tissue engineering and the contemporary use of stem cells in the field of urological research. It will also provide a review of leading stem cell procurement techniques utilized in contemporary tissue engineering labs. These include embryonic stem cell harvest, parthenogenesis, nuclear transfer, and most recently, induced pluripotent stem cells (iPS cells). It is our intention to leave the reader with a solid foundation of knowledge concerning the use of stem cells in regenerative medicine and its present and future applications in urology.

17.2

Principles of tissue engineering

Current practices in reconstructive surgery utilize autologous tissue to avoid histocompatibility issues and possible rejection. When autologous

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tissue is lacking, other tissue sources include homologous tissues from cadavers or donors, heterologous tissues from animals (bovine), and artificial materials (silicone, polyurethane, Teflon) often referred to as ‘alloplastic’. There are numerous synthetic materials available for prosthetics and other clinical devices that are biocompatible and provide structural integrity. However, functional recovery to the caliber of the original tissue is seldom achieved with synthetic substitutes. Use of allogenic tissues, i.e. organ transplantation, is limited by the need for tissue matching, the adverse effects of lifelong immunosuppression, and donor availability. Tissue engineering offers the potential to circumvent many of these obstacles. Tissue engineering approaches can be classified into two broad categories: acellular and cellular techniques. Acellular techniques use biological matrices devoid of cells as scaffolds for organ regeneration, requiring the host organ to generate new tissue into the scaffold with proper layering and orientation. These acellular scaffolds, which are commonly composed of collagen-rich matrices, can be completely synthesized in the laboratory. Their substrates are harvested from autologous, allogenic, or xenogenic tissues, and then processed by chemical and mechanical means to remove cellular components for eventual implantation (Chen et al., 1999). Once implanted into host tissue, these matrices slowly degrade and are eventually replaced by host extracellular matrix produced by in-growing cells. Cellular techniques utilize cells that are acquired from xenogenic, allogenic, or autologous sources and also typically processed prior to implantation. These cells can be applied via direct injection or seeded on to a support matrix or scaffold, followed by implantation into the recipient. As previously stated, donor tissue/cell sources vary, with autologous cells being the most preferred to eliminate the risk of rejection and associated complications of immunosuppression. For autologous sources, a tissue sample can be obtained from a patient by biopsy (Atala et al., 2006; Yoo et al., 1998). Improvements in culture techniques have enabled the isolation of individual cell types from these tissue biopsies, which are then selectively expanded to amounts sufficient for implantation (Amiel et al., 1999; Cilento et al., 1994). Scaffold materials must be biocompatible, bio-resorbable, and display minimal immunogenicity to allow expanding cellular components to integrate successfully into the local environment. Scaffolds may be seeded or impregnated with growth factors and other cell signaling peptides to regulate cell activity and mimic the microenvironment provided by the extracellular matrix. Ideal scaffolds should also provide an appropriate three-dimensional lattice where cell adhesion may occur while performing the mechanical functions of the damaged tissue, thus allowing the final steps of the regenerative process to occur in vivo (Kim and Mooney, 1998).

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Techniques utilizing direct injection of stem cells do not require the use of tissue scaffolds. Stem cells derived from autologous tissues and expanded in culture are implanted directly into the organ or structure in need of repair, where their replication and differentiation occurs in the host extracellular environment. While these techniques may not be efficacious in the development of complex organ structures, their comparative simplicity makes them an attractive option for urological applications where simple tissue growth is all that is required.

17.3

Stem cells: overview

It is primarily in the field of cellular tissue engineering techniques that stem cell research has the potential to provide important advances. Tissue grafts created from stem cells could provide prefabricated ‘off the shelf’ cellular tissue replacements. In engineering non-immunogenic reproducible stem cells, one could avoid the additional step of tissue biopsy, and produce a readily available source of non-patient-specific tissue grafts, as well as cells for injection therapy. Stem cells are defined by their ability to self-renew and differentiate into a wide array of mature cell types (Becker and Jakse, 2007). Stem cells are inherently difficult to classify because of their obscure morphological and molecular characteristics. Rather, they are classified according to their potency. Cells that possess the potential to differentiate into the widest array of cell types are termed ‘totipotent’. Examples of these include the zygote and offspring cells of the morula, which can differentiate into endoderm, ectoderm, mesoderm, cells of the gonadal ridge, and ‘extraembryonic’ cells that support the embryo. Human embryonic stem cells isolated from the blastocyst are termed ‘pluripotent’ and can give rise to all cell types except for extra-embryonic cells. ‘Multipotent’ stem cells are capable of self-renewal, but are often limited in their ability to differentiate into cell types of other specific organs. ‘Unipotent’ stem cells exhibit limited or no capacity for self-renewal and differentiate into only one cell type (Table 17.1). In the search for efficiently reproducible tissue substrates, various stem cell sources have been examined. Of the various available sources, human embryonic stem cells (hESCs) have been shown to provide the stem cell source of highest differentiation potential. According to data from the Centers for Disease Control, it has been estimated that approximately 3000 Americans die every day of diseases that could have been treated with embryonic stem cell-derived tissues (Lanza et al., 2001). However, their procurement has been characterized by widespread controversy because it requires the destruction of human embryos. Thus, other forms of stem cell production have been investigated in an effort to circumvent the ethical

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Table 17.1 Embryonic stem cells Embryonic stem cell type Totipotent Pluripotent

Multipotent

Unipotent

Definition

Sources/examples

Able to differentiate into all cell types Able to differentiate into all cell types except extra-embryonic cells Capable of self-renewal, limited ability to differentiate into cell types of other organ systems Limited or no capacity for self-renewal, differentiate into one cell type

Zygote, offspring cells of the morula Blastocyst

Adult stem cells: hematopoietic stem cells, spermatogonia, gastrointestinal crypt cells Most adult somatic cells

Table 17.2 Stem cell sources Method

Definition

Disadvantages

hESCs

Harvest from blastocyst of embryo

Nuclear transfer

Transfer of somatic cell nucleus into enucleated oocyte (cloning) Stimulation of unfertilized oocyte into embryogenesis ‘Reprogramming’ of adult somatic cells into pluripotent stem cells

Ethical concerns over destruction of human embryos Ethical concerns over reproductive cloning Requires ovarian biopsy for oocyte procurement Relatively low yield of pluripotent cells with current reprogramming protocols

Parthenogenesis Induced pluripotent cells

disputes associated with hESCS. These include parthenogenesis, nuclear transfer, and iPS cells (Table 17.2).

17.4

Embryonic stem cells

17.4.1 Overview hESC research holds enormous therapeutic potential. Proposed uses include regeneration of nervous tissue in the treatment of spinal cord injury, re-growth of healthy kidney parenchyma in patients with renal failure, and cure of insulin-dependent diabetes via implantation of healthy pancreatic

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islet cells. These are but a few examples of the manner in which stem cells could transform chronic debilitating illnesses into curable conditions, thus justifying the magnitude of interest of both the scientific community and the popular media in the development of this technology. Embryonic stem cells exhibit two remarkable properties: the ability to proliferate in an undifferentiated state, and the ability to differentiate into many specialized cell types (Brivanlou et al., 2003). Although they give rise to all three germ layers (endoderm, mesoderm, ectoderm), they do not give rise to extra-embryonic tissues and are thus termed pluripotent. hESCs can be isolated with immunosurgery from the inner cell mass of the embryo during the blastocyst stage (5 days post-fertilization), and are usually grown on feeder layers consisting of mouse embryonic fibroblasts or human feeder cells (Richards et al., 2002). More recent reports have shown that hESCs can be grown without the use of a feeder layer (Amit et al., 2004), which avoids the exposure of these cells to murine viruses and proteins. These cells have demonstrated longevity in culture by maintaining their undifferentiated state for at least 80 passages when grown using current published protocols (Reubinoff et al., 2000; Thomson et al., 1998). Differentiation of hESCs into cells from all three embryonic germ layers has been demonstrated in vitro. Skin and neurons have been formed, indicating ectodermal differentiation (Reubinoff et al., 2001; Schuldiner et al., 2000; Schuldiner et al., 2001; Zhang et al., 2001). Blood, cardiac cells, cartilage, endothelial cells, and muscle have been formed, indicating mesodermal differentiation (Kaufman et al., 2001; Kehat et al., 2001; Levenberg et al., 2002). Pancreatic cells have been formed, indicating endodermal differentiation (Assady et al., 2001). In addition, as further evidence of their pluripotency, embryonic stem cells can form embryoid bodies in vitro – which are cell aggregations that contain all three embryonic germ layers – and can form teratomas in vivo (Itskovitz-Eldor et al., 2000).

17.4.2 Clinical applications In the field of urology, there is ongoing research into the use of hESCs to develop bladder substitutes. Severe bladder diseases, such as invasive bladder carcinoma and severe persistent hemorrhagic cystitis, require surgical removal of the bladder, thus precluding the use of autologous cells obtained via bladder biopsy for tissue regeneration. In these patients, the use of hESCs as a primary non-immunogenic tissue source has the potential to be therapeutically invaluable. Successful directed differentiation of human embryoid body-derived stem cells into bladder urothelium has been reported in a rat model. This was accomplished by co-culture of the stem cells with bladder mesenchyme, thus providing the stimulatory effects of mesenchymal inductive co-factors

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(Lakshmanan et al., 2005). The seeding of these stem cell/ mesenchyme cocultures on decellularized xenogenic small intestine submucosa has been reported to produce composite bladder urothelial grafts successfully. Although the function of these grafts was not tested in vivo, these studies demonstrate the feasibility of generating stem cell-derived bladder substitutes (Frimberger et al., 2005). Several studies have focused on the use of hESCs in the formation of prostatic glandular tissue. Although the majority of prostatic disease presents at a later stage in life, and appropriate therapy centers on the removal of hypertrophied or cancerous tissue rather than on preservation of reproductive function, the findings of these studies represent important advances in the use of stem cells for the regeneration of urological tissues and organs. One recent study described in vivo directed differentiation of hESCs into prostatic tissue using tissue recombination techniques (Taylor et al., 2006). Heterospecific tissue recombinants were created with hESCs and rat and murine urogenital sinus mesenchyme. The recombinant tissue was then implanted into the renal cortex of immunodeficient mice. Findings included the achievement of glandular differentiation, which was further demonstrated by the production of prostate-specific antigen. Interestingly, controls grown in the absence of mesenchyme produced teratomas.

17.4.3 Ethical issues The harvest of hESCs requires the destruction of human embryos and has thus raised significant ethical and political concerns. In August 2001, the United States federal government ordered that only previously generated hESC lines could be used in research supported by federal funding (Vogel, 2001). Although over 70 existing cell lines met this criterion, the National Institutes of Health classified only 11 as available for research, most of which were grown on mouse feeder cells and were at one point in time potentially exposed to murine viruses or proteins (Kennedy, 2003; Wertz, 2002). This significant restriction on the development of hESC technology has prompted research into alternative stem cell procurement techniques and sources. These include the use of fetal tissues, parthenogenesis, amniotic fluid-derived stem cells, somatic cell nuclear transfer, and adult multipotent stem cells.

17.5

Nuclear transfer

17.5.1 Overview Nuclear transfer, which has also been called nuclear transplantation and nuclear cloning, involves the introduction of a nucleus from a donor cell into a previously enucleated oocyte, thus generating an embryo with a

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genetic make-up identical to that of the donor. While there has been tremendous interest in the field of nuclear cloning since the birth of Dolly in 1997, the first successful nuclear transfer was reported over 50 years ago by Briggs and King (1952). Cloned frogs, which were the first vertebrates derived from nuclear transfer, were subsequently reported by Gurdon (1962), but the nuclei were derived from non-adult sources. Dolly was not the first cloned mammal to be produced from adult cells; live lambs were produced in 1996 using nuclear transfer and differentiated epithelial cells derived from embryonic discs (Campbell et al., 1996). Despite these early successes, the field has remained relatively undeveloped, and only in the last decade have many of the most important advances in nuclear cloning technology occurred. The significance of Dolly was that she was the first mammal to be derived from an adult somatic cell using nuclear transfer (Wilmut et al., 1997). Since then, animals from several species have been grown using nuclear transfer technology, including cows (Cibelli et al., 1998), goats (Baguisi et al., 1999; Keefer et al., 2002), mice (Wakayama et al., 1998), and pigs (Betthauser et al., 2000; De Sousa et al., 2002; Onishi et al., 2000; Polejaeva et al., 2000).

17.5.2 Reproductive cloning Two types of nuclear cloning (reproductive cloning and therapeutic cloning) have been described, and a better understanding of the differences between the two types may serve to better elucidate some of the controversy that surrounds these technologies (Colman and Kind, 2000; Vogelstein et al., 2002). Banned in most countries for human application, reproductive cloning is used to generate an embryo whose genetic material is identical to that of its cell donor. This embryo is then implanted into the uterus of a female to give rise to an infant that is a clone of the donor. Alternatively, in therapeutic cloning the embryo is explanted into a tissue culture medium to produce embryonic stem cell lines whose genetic material is identical to that of its source. These cultured autologous stem cells have the potential to become almost any type of cell in the adult body, and thus would be invaluable in tissue and organ replacement (Hochedlinger and Jaenisch, 2003). Potential clinical applications include the treatment of end-stage kidney disease, neurodegenerative diseases, and diabetes, conditions for which there is limited availability of immunocompatible tissue transplants.

17.5.3 Clinical applications Stem cells harvested from in vitro fertilized human embryos, hESCs, are genetically divergent from the patient, therefore necessitating immunosuppression to prevent rejection by their host. Indeed, with current allogeneic

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tissue transplantation protocols, rejection is a frequent complication because of immunological incompatibility, and immunosuppressive drugs are usually administered to prevent and treat host-versus-graft disease (Hochedlinger and Jaenisch, 2003). The use of transplantable tissues and organs derived from therapeutic cloning may lead to the avoidance of this immune response, as these tissues are produced from stem cells that are genetically identical to their hosts. Therefore, with therapeutic cloning, the variety of serious and potentially life-threatening complications associated with immunosuppressive treatments may be avoided (Lanza et al., 1999). The application of nuclear transfer technology to urological reconstruction was demonstrated by Lanza and Atala (Lanza et al., 2002) when they successfully engineered kidney tissue genetically identical to a host bovine model. Unfertilized donor oocytes were harvested and the nuclei removed 18–22 hours after maturation. Skin fibroblasts from adult Holstein steers were obtained by ear notch, and single donor cells microinjected into the perivitelline space of the enucleated oocytes, thus completing nuclear transfer. An electrical pulse was administered to initiate fusion of the cell–oocyte complexes resulting in the formation of blastocysts. These were implanted into recipient host bovines for 12 weeks of in vivo growth. Cloned renal cells were then harvested from the blastocysts, expanded in vitro, and seeded on to specially designed biodegradable scaffolds. These scaffolds consisted of three collagen-coated cylindrical polycarbonate membranes, the ends of which were connected to catheters that terminated into a collecting reservoir. The cell–scaffold constructs, which constitute a renal neoorgan with a mechanism for collecting excreted fluid, were then implanted into the subcutaneous space of the original donor steer to allow for tissue growth for 12 additional weeks. The collected fluid from the constructs, which resembled urine in appearance, underwent chemical analysis. Urea nitrogen and creatinine were observed to be near physiological urine levels, and measurement of electrolyte levels, specific gravity, and glucose concentration, revealed that the implanted renal units demonstrated reabsorption, filtration, and secretory capabilities. Histological examination of the explanted units revealed extensive vascularization and self-organization of the cells into structures resembling glomeruli and tubules, with a geometric arrangement that allowed the passage of urine into the collecting reservoir. Immunohistochemical and polymerase chain reaction (PCR) analyses were also performed, revealing the presence of renal-specific proteins and transcription of renal cell-specific RNA. Identified proteins included synaptopodin (produced by podocytes), aquaporin-1 (produced by proximal tubules and the descending limb of the loop of Henle), aquaporin-2 (produced by the ascending loop of Henle), Tamm–Horsfall protein (produced by the ascending loop of Henle), and Factor VII (produced by endothelial

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cells). Two physiologically important renal protein products, erythropoietin and 1,25-dihydroxyvitamin D3, were also present. This study also investigated the concern that the donor egg’s mitochondrial DNA may be a source of immunological incompatibility (Evans et al., 1999; Hiendleder et al., 1999; Steinborn et al., 2000). This could stimulate a T-cell response for mt-DNA-encoded minor histocompatibility antigens upon implantation of the cloned cells into the original nuclear donor (Fischer Lindahl et al., 1991). Maternally transmitted minor histocompatibility antigens in mice have been shown to stimulate both skin allograft rejection in vivo and cytotoxic T-lymphocyte expansion in vitro (Fischer Lindahl et al., 1991), which could preclude the usage of cloned renal constructs in patients with chronic rejection of major histocompatibilty matched human renal transplants (Hadley et al., 1992; Yard et al., 1993). Testing for T-cell response to the cloned renal units was conducted with delayed-type hypersensitivity testing in vivo and Elispon anlysis of interferon-gammasecreting T-cells in vitro. Both tests demonstrated no evidence of a T-cell response, suggesting that the presence oocyte-derived mitochondrial DNA (mt-DNA) does not necessarily elicit rejection. This finding may represent a step forward in overcoming the histocompatibility problem of stem cell therapy (Auchincloss and Bonventre, 2002). The kidney is a complex organ with multiple cell types and a complex functional anatomy that renders it one of the most difficult to reconstruct (Amiel and Atala, 1999; Auchincloss and Bonventre, 2002). Previous efforts in tissue engineering of the kidney have been directed toward the development of extracorporeal renal support systems made of biological and synthetic components (Aebischer et al., 1987; Amiel et al., 2000; Humes et al., 1999; Ip et al., 1988; Joki et al., 2001; Lanza et al., 1996; MacKay et al., 1998), and ex vivo renal replacement devices (hemodialysis) are known to be lifesustaining. However, there would be obvious benefits for patients with end-stage kidney disease if these devices could be implanted long-term without the need for an extracorporeal perfusion circuit or immunosuppressive drugs. The demonstrated ability of using stem cells harvested via nuclear transfer to produce implantable functioning renal units that are potentially non-immunogenic clearly represents a quantum leap forward in the field of renal replacement technology. This study demonstrated that cells derived from nuclear transfer can be successfully harvested, expanded in culture, and transplanted in vivo with the use of biodegradable scaffolds on which the single suspended cells can organize into tissue structures that are genetically identical to those of the host. It represents substantial progress toward the eventual goal of being able use a single skin cell from a patient to generate any type of tissue that could be replaced or transplanted. An advantage of this system, as opposed to current transplantation techniques, is that the transplanted tissue would

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be genetically identical to that of the recipient and would be fully immunocompatible. Conceivably, each patient could have a ready-made supply of their tissues available for their own use, should the need arise in the future.

17.5.4 Ethical issues The ethical controversies surrounding nuclear transfer technology have thus far centered on use of the technique to create living clones, e.g. reproductive cloning. Nuclear transfer for therapeutic cloning pre-empts this concern, as the embryonic tissues derived are not implanted into a uterus for development into a viable living organism. Stem cells produced by nuclear transfer do not necessitate the destruction of a viable human embryo, and are thus less prone to the ethical and moral dilemmas that have impeded the development of hESC technology.

17.6

Parthenogenesis

17.6.1 Overview Parthenogenesis (Greek: ‘virgin birth’) is a form of reproduction in which an unfertilized egg develops into a new individual. Parthenogenesis occurs naturally in certain plants, arthropods, and insects. Artificial parthenogenesis with frog eggs was first described by Loeb in 1900, and first described in mammalian eggs by Pincus in 1936. Parthenogenesis represents an important alternative source of pluripotent embryonic stem cells in that it establishes the development of an embryo without fertilization by sperm, and without the transfer of nuclear material into an ovum (Cowan et al., 2004). As evidence of the multipotency of these cells, Kuno et al. (2004) noted the presence of fully differentiated tissue of all three germ layers in a mature ovarian cystic teratoma of a virginal woman, thereby suggesting that the information for the organization of the body plan is already present in the oocyte. Recently, parthenogenesis with non-human primate eggs has been identified as a potential source for embryonic stem cells (Cibelli et al., 2002); however, the cells were differentiated only to the neuronal lineage, and demonstration of multipotency was limited to teratoma formation with the presence of all three germ layers.

17.6.2 Clinical applications As noted above, previous studies have described the isolation of embryonic stem cells from parthenogenesis (Brevini and Gandolfi, 2008; Cheng, 2008; Cibelli et al., 2002; Dighe et al., 2008; Fang et al., 2006; Lin et al., 2007; Mai

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et al., 2007; Sritanaudomchai et al., 2007). In addition, we have recently described the adaptation of parthenogenetic stem cells for tissue engineering applications, which includes the growth, harvest, and culturing of stem cells obtained via parthenogenesis into multiple tissue lines that potentially may be used for future therapeutic purposes (Koh et al., 2009). Oocytes in meiosis II collected from superovulated New Zealand rabbits were activated with electrical stimulation, then incubated for 2 hours. Once the activated oocytes reached blastocyst stage, microsurgery was used to dissect the inner cell mass from the blastocyst. This inner cell mass was expanded in culture. Multipotent stem cells, which constituted approximately 10% of the total cell population, were then immunoisolated from the rest of the cell population using stem cell markers and expanded with a doubling time of approximately 20 hours. Cell cycle analysis and karyotyping confirmed these stem cells to be homogeneously diploid. Further analysis demonstrated the presence of several important early embryonic stem cell markers in these cells, reflecting their potential to develop into multiple tissue types. The cells were then induced to osteogenic (Jaiswal et al., 1997), myogenic (Ferrari et al., 1998; Rosenblatt et al., 1995), and adipogenic (Poliard et al., 1995) lineages by manipulation of their specific growth conditions as described by standardized protocols. Cells were also induced to an endothelial lineage within a specially supplemented endothelial cell medium. Their differentiation was confirmed by phenotypic changes, immunocytochemistry, gene expression, and functional analyses. Cells directed toward the myogenic lineage demonstrated the presence of desmin, myoD, actinin, and sarcomeric tropomysin, all early muscle markers. The osteogenically directed cells demonstrated the presence of calcium mineralization, osteocalcin, which is synthesized singularly by osteoblasts, and osteopontin, another important bone-specific protein. Cells directed toward the adipogenic lineage contained lipid droplets, and demonstrated expression of adipogenic genes such as lipoprotein lipase. Cells differentiated towards endothelium demonstrated a cobblestone appearance on microscopy and several molecular markers specific for endothelial cells. The differentiated stem cells were then tested in vivo. The myogenic and osteogenic cell lines were labeled and then seeded on to acellular scaffolds and implanted into athymic mice for 2 and 4 weeks, respectively. After reharvest, labeled muscle-like and bony-like tissue was obtained. A murine muscle injury model was then created by injection of cardiotoxin into the tibialis muscle in mice. The myogenic stem cells were injected into the injured muscle, and 2 weeks later, histopathological evaluation revealed the survival and integration of these stem cells into the injured muscle. This study is remarkable in that it demonstrated the successful adaptation of parthenogenesis-derived stem cells for tissue engineering, and that stem cells derived in this manner showed a rapid proliferation rate that would

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allow adequate expansion of these cells for tissue replacement therapy. Also encouraging is that the source of these stem cells, oocytes, can be harvested with minimally invasive techniques and low donor morbidity. The methods described in this experiment will hopefully provide a useful framework for the eventual production of large quantities of human cells and tissues for regenerative medicine applications.

17.6.3 Ethical issues The use of parthenogenesis-derived stem cells may avoid some of the political and ethical controversy surrounding hESCs. Several attempts to produce viable individuals from parthenogenesis in the past have been unsuccessful (Fukui et al., 1992; Kaufman, 1977; McGrath and Solter, 1984). While Kono and colleagues (2004) have recently shown that some parthenogenetic mice can survive into adulthood, 99.4% of the embryos perished prior to adulthood, with the majority of the deaths occurring during embryonic development. This may help to relieve some of the potential ethical and political concerns with the use of parthenogenetic stem cells in that the parthenogenetic embryos rarely survive beyond the embryo stage, but do survive long enough for the embryonic stem cells to be harvested for tissue engineering purposes. The ability to procure embryonic stem cells without sacrificing human embryos created through conception thus represents an invaluable resource for patients who require the replacement of damaged or diseased tissue.

17.7

Induced pluripotent stem cells

iPS cells represent a relatively new and potentially very significant source of stem cells. Indeed, a large proportion of current stem cell and tissue engineering research is focused on developing this technology. iPS cells are produced by transformation of adult somatic cells into an embryonic stem cell-like state through injection of specific genes via viral vectors, a process known as direct reprogramming (Mikkelsen et al., 2008). A protocol using four genes, including the cancer cell-related c-myc oncogene, was first reported by Takahashi and Yamanaka (2006) with successful reprogramming of mouse fibroblasts into iPS cells. Since then, several groups of researchers have reported generation of iPS cells from human tissues using similar protocols (Meissner et al., 2007; Park et al., 2008; Takahashi et al., 2007; Yu et al., 2007). Others have reported successful reprogramming without use of the c-myc oncogene (Wernig et al., 2008). These studies demonstrated the generated human iPS cells to be similar to hESCs in morphology, proliferation, surface antigens, and gene expression, among other similarities. Human iPS cells have also been shown to possess the

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ability to differentiate into cell types of all three germ layers in vitro as embryoid bodies and in vivo as teratomas (Yu et al., 2007). The mechanism of this dramatic transformation centers on the observation that only a small subset of a cell’s gene library is expressed at any given time, and that the expressed subset varies by cell type (Costello, 2008). Gene expression is controlled by many cellular processes, including DNA methylation, chemical modification of chromatin, and the activities of transcription factors. These processes, collectively referred to as ‘epigenetics’, control which genes are available for transcription and the rates at which they are transcribed, ultimately affecting the phenotype of the cell. iPS cell technology allows the use of readily available adult somatic cells to create stem cells, thus circumventing many of the ethical implications of utilizing fetal and embryonic tissues (Costello, 2008). The production of iPS cells also avoids the technical difficulties and donor risks associated with the harvest of oocytes, which is necessary in somatic cell nuclear transfer. The introduction of iPS cell technology has spurred many to declare that the harvest of stem cells from human embryos is no longer necessary and should thus be abandoned completely. Many leading researchers believe this sentiment to be premature and unrealistic. This is because it remains to be proven whether iPS cells are in fact equivalent to embryonic stem cells, both in their make-up and in their therapeutic potential. Simply determining whether they are equivalent will require embryonic stem cells for comparison, and the current need to inject retroviruses into iPS cells as part of the reprogramming process raises significant safety concerns for human implantation (Holden and Vogel, 2008).

17.8

Conclusion and future trends

The field of regenerative medicine has made many important advances recently in the area of tissue engineering, as evidenced by the multitude of successful applications of tissues grown in culture for urological and other medical therapies. The majority of successful clinical applications have thus far focused on the expansion of autologous cells in culture with subsequent implantation (De Filippo et al., 2002; De Filippo et al., 2003; Oberpenning et al., 1999). This technique has several advantages over the use of allogenic transplantation (i.e. kidney transplant) or substitution of other autologous tissues (i.e. ileal bladder augmentation), including non-immunogenicity, increased availability, and improved physiological function. The use of engineered bladder tissues for surgical augmentation in patients with neurogenic bladders and poor bladder capacity is a current example of one exciting therapeutic application of this technology (Atala et al., 2006). Engineering autologous somatic cell cultures for tissue replacement is currently complicated in that it requires significant resources for successful

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application. Its efficiency is diminished by the need for potentially risky biopsy procedures, and the time required for growth of cells in culture. Furthermore, many patients who require replacement tissues may not have a satisfactory autologous source (i.e. cancer patients). Stem cells present a potentially superior source of tissues for regenerative medicine, in that stem cells have the potential to overcome many of these shortcomings. Stem cells can be induced to grow into multiple tissue types, can be stored and expanded in sustainable cell lines, and thus be used to develop a stock of readily available cells for replacement of virtually any tissue. The initial excitement over the potential curative applications of hESCs has been tempered by ethical concerns over the destruction of human embryos. Although hESCs have been the most studied type of stem cell, and are considered to be the gold standard in many respects, other methods of stem cell harvest are currently under development that do not necessitate destruction of a human embryo, and which could therefore circumvent many of the ethical dilemmas impeding stem cell research. Some of these methods, such as stem cell harvest from induced pluripotent cells, may also prove to be far more efficient than embryonic stem cell harvest. The future directions of tissue engineering and regenerative medicine will likely focus largely on the use of stem cells for a broad range of therapeutic applications. The search continues for new, readily available sources of stem cells, some of which have been described above. The recent interest in iPS cells will likely intensify, as researchers investigate methods to improve the efficiency of their manufacture and search for new cell reprogramming protocols that do not require the use of potentially harmful viruses. Future research with iPS cells will also likely focus on the use of these cells, easily harvested from an individual patient’s skin, to test patientspecific responses to current and future medical therapies (Holden and Vogel, 2008). Amniotic fluid, which contains many fetal cells in suspension, has been investigated as a source of pluripotent stem cells. De Coppi and colleagues (2007) were able to demonstrate induction of human amniotic fluid stem cells toward specific lineages and implantation of these cells in vivo to generate specialized tissue, include neural tissue and bone, in a mouse model. Their demonstrated ability to produce functioning neural tissues in vivo from stem cells underscores the immense potential for stem cell technology to cure many of humanity’s most debilitating diseases.

17.9

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18 Amniotic fluid and placental stem cells as a source for urological regenerative medicine P. D E C O P P I, UCL Institute of Child Health and Great Ormond Street Hospital, UK and Azienda Ospedaliera Università di Padova, Italy; G. BA RT S C H, University of Ulm, Germany; and A. ATA L A, Wake Forest Institute for Regenerative Medicine, USA

Abstract: Stem cells capable of differentiating to multiple lineages may be valuable for urological regenerative medicine. Recent research has described the presence in the amniotic fluid of stem cells with various differentiative and proliferative potentials. The use of amniotic fluidderived stem cells does not present the same ethical challenges as the use of embryonic stem cells and the cells can be isolated using amniocentesis, a widely accepted procedure for prenatal diagnosis. This chapter summarizes the different progenitor cells that have been described so far in the amniotic fluid and the placenta and focuses in particular on mesenchymal stem cells and amniotic fluid-derived stem cells. Key words: amniotic fluid-derived stem cells, regenerative medicine.

18.1

Introduction

In recent years the presence in the amniotic fluid of stem cells with various differentiative and proliferative potentials has been described. Amniocentesis is a widely accepted method for prenatal diagnosis. Minimal ethical concerns arise if fetal stem cells are isolated from amniotic fluid prior to or at the time of birth. In this chapter we will briefly describe the technique of amniocentesis and the different progenitor cells described to date.

18.2

Amniocentesis

One of the primary applications of amniocentesis is as a safe method of isolating fetal cells for chromosomal analyses (Milunsky, 1979; Hoehn and Salk 1982; Gosden 1983; Crane and Cheung, 1988). The protocol generally consists of acquiring 10–20 ml of amniotic fluid using a transabdominal approach. Amniotic fluid samples are centrifuged, and the cell pellet is 378

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resuspended in culture medium. Approximately 104 cells are seeded on 22 mm × 22 mm cover slips. Cultures are grown to confluence for 3–4 weeks in 5% CO2 at 37 °C, and the chromosomes are characterized from mitotic phase cells (Brace and Resnik, 1999). Amniocentesis is usually performed around the 16th week of gestation, although in some cases it may be performed as early as the 14th week of gestation when the amniotic membrane fuses with the chorionic membrane and the risk of bursting the amniotic sac by needle puncture is minimized. Amniocentesis may be performed as late as term. The amniotic sac is usually noticed first by ultrasound around the 10 week gestational time point.

18.3

Differentiated cells from amniotic fluid

Amniotic fluid cell culture consists of a heterogeneous cell population displaying a range of morphologies and behaviors. Phenotypic studies have characterized these cells into many shapes and sizes varying from 6 μm to 50 μm in diameter and from a round to a squamous cell shape. Most cells in the fluid are terminally differentiated along epithelial lineages and have limited proliferative and differentiation capabilities. Previous studies have noted an interesting composition of the fluid consisting of a heterogeneous cell population expressing markers from all three germ layers (Sarkar et al., 1980; Cousineau et al., 1982; Medina-Gomez and Johnston, 1982; von Koskull et al., 1984). Current theories suggest that the fluid is mainly derived from urine and pulmonary secretion of the fetus as well as from some ultrafiltrate of the mother’s plasma entering through the placenta. The cells in the fluid have been shown to be overwhelmingly from the fetus, mostly cells sloughed off the skin and from the digestive and urinary tract of the fetus, as well as from the amniotic membrane (Lotgering and Wallenburg, 1986; Underwood et al., 2005).

18.4

Mesenchymal stem cells from amniotic fluid

Several studies have been published in the last few years describing simple protocols for the isolation from amniotic fluid of mesenchymal stem cells (MSCs) similar to those that have been described as being present in various adult tissues such as bone marrow (Haigh et al., 1999; Kaviani et al., 2001; Kaviani et al., 2003). These cells were able to proliferate in vitro, to be engineered in a three-dimensional structure and used in vivo to repair tissue defects (Kaviani et al., 2003). A few years later In’t Anker et al. (2003) were able to prove for the first time that both amniotic fluid and placenta were abundant sources for fetal MSCs that exhibit a phenotype and multilineage differentiation potential similar to that of postnatal bone marrowderived MSCs (In’t Anker et al., 2003). Isolation and expansion protocols

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are similar to those used to isolate MSCs from other sources. Briefly, amniotic fluid samples were centrifuged for 10 minutes at 1283 rpm. Pellets were resuspended in Iscove’s modified Dulbecco’s medium containing 2% fetal calf serum (FCS) and antibiotics (defined as washing medium). Similarly, for the placenta, approximately 1 cm2 was washed in phosphate-buffered saline (PBS) and single-cell suspensions were made by mincing and flushing the tissue parts through a 100 μm nylon filter with washing medium. Singlecell suspensions of amniotic fluid and placenta tissue were plated in six-well plates and cultured in M199 supplemented with 10% FCS, 20 μg/ml endothelial cell growth factor, heparin (8 U/ml) and antibiotics. After 7 days, non-adherent cells were removed and the medium was refreshed. When grown to confluence, adherent cells were detached with trypsin ethylene diaminetetraacetic acid (EDTA) and expanded in culture flasks pre-coated with 1% gelatin and kept in a humidified atmosphere at 37 °C. The expansion potency of fetal MSCs was higher compared with that of adult bone marrow-derived MSCs. As a result, fetal MSCs were able to expand amniotic fluid MSCs to about 180 × 106 cells within 4 weeks (three passages). The phenotype of the culture-expanded amniotic fluid-derived cells was similar to that reported for MSCs derived from second-trimester fetal tissues and adult bone marrow. In’t Anker et al. were able to show that amniotic fluid-derived MSCs showed the potential for multilineage differentiation into fibroblasts, adipocytes and osteocytes (In’t Anker et al., 2004). Furthermore, amniotic fluid-derived MSCs were successfully isolated, cultured and enriched without interfering with the routine process of fetal karyotyping. Flow cytometry analyses showed that they were positive for SH2, SH3, SH4, CD29 and CD44, weakly positive for CD90 and CD105, but negative for CD10, CD11b, CD14, CD34, CD117 and EMA (Tsai et al., 2004). Most importantly, immunophenotypic analyses demonstrated that these cells expressed HLA-ABC (class I major histocompatibility complex (MHC-I)), but they did not express HLA-DR, -DP, -DQ (MHC-II molecules) (Li et al., 2005). Li et al. (2005) have investigated the immunological role of amniotic fluid-derived MSCs extensively. They have reported that mononucleated cells recovered from placentas by density gradient fractionation could suppress umbilical cord blood (UCB) lymphocyte proliferation induced by cellular or non-specific mitogenic stimuli. This immunoregulatory feature strongly implies that the amniotic fluid-derived MSCs may have potential applications in allograft transplantation. Since it is possible to obtain placenta and UCB from the same donor, they suggested the placenta as an attractive source of MSCs for co-transplantation in conjunction with UCB-derived hematopoietic stem cells (HSCs) to reduce the potential graft-versus-host disease (GVHD) in recipients (Li et al., 2005). Finally, ovine mesenchymal amniocytes have also been cultured and engineered into a collagen hydrogel in order to replace partial diaphragmatic

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losses or absences (Fuchs et al., 2004). The authors showed that diaphragmatic repair with an autologous tendon engineered from mesenchymal amniocytes leads to improved mechanical and functional outcomes when compared with an equivalent acellular bioprosthetic repair. Different groups have shown that MSCs from placenta and amniotic fluid may have more plasticity than was initially thought. Phenotypic and gene expression studies indicated mesenchymal stem cell-like profiles in both amnion-derived and chorion-derived cells that were positive for neuronal, pulmonary, adhesion and migration markers. In addition, transplantation in neonatal swine and rats resulted in human microchimerism in various organs and tissues, suggesting that amnion and chorion cells may represent an advantageous source of progenitor cells with potential applications in a variety of cell therapy and transplantation procedures (Bailo et al., 2004). Similarly, Zhao et al. (2005) have reported that human amniotic mesenchymal cells (hAMCs), may also be a suitable cell source for cardiomyocytes. They showed that freshly isolated hAMCs expressed cardiac-specific transcription factor GATA4, cardiac-specific genes – such as myosin light chain (MLC)-2a, MLC-2v, cTnI, and cTnT – and the alpha-subunits of the cardiac-specific L-type calcium channel (alpha1c). After stimulation with basic fibroblast growth factor (bFGF) or activin A, hAMCs expressed Nkx2.5, a specific transcription factor for the cardiomyocyte and cardiacspecific marker atrial natriuretic peptide. In addition, the cardiac-specific gene alpha-myosin heavy chain was detected after treatment with activin A. Co-culture experiments confirmed that hAMCs were able to both integrate into cardiac tissues and differentiate into cardiomyocyte-like cells. After transplantation in rat hearts after acute myocardial infarcts, hAMCs survived in the scar tissue for at least 2 months and differentiated into cardiomyocyte-like cells (Zhao et al., 2005). However, we have recently shown that this potential is not exhibited by mesenchymal progenitor cells in larger mammals such as pigs. Amniotic fluid-derived mesenchymal cells autotransplanted in a porcine model of acute myocardial infarcts were able to transdifferentiate to cells of vascular cell lineages but failed to give origin to cardiomyocytes (Sankar and Muthusamy, 2003; Sartore et al., 2005). Neuron regeneration has been also described using rat amniotic epithelial cells. In particular they expressed in vitro both neuronal and neural stem cell markers, neurofilament microtubule-associated protein 2 and nestin. Reverse-transcriptase-polymerase chain reaction (RT-PCR) revealed that these cells expressed nestin. The rat amniotic epithelial cells were also transplanted into the hippocampus of adult gerbils that were subjected to temporal occlusion of bilateral carotid arteries. Five weeks after transplantation, grafted cells migrated into the CA1 pyramidal layer which showed selective neuronal death, and survived in a manner similar to CA1 pyramidal neurons (Okawa et al., 2001). Different reports suggest that human

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amniotic epithelial cells (HAECs) also possess certain properties similar to those of neural and glial cells (Tsai et al., 2006). When transplanted into the transection cavities in the spinal cord of bonnet monkeys, HAECs were able to survive, support the growth of host axons through them and prevent glial scar formation at the cut ends, and may prevent death in axotomized cells or attract the growth of new collateral sprouting (Okawa et al., 2001). AECs isolated from human term placenta express surface markers normally present on embryonic stem and germ cells. In addition, they express the pluripotent stem cell-specific transcription factors octamer-binding protein 4 (Oct-4) and nanog. Under certain culture conditions, AECs form spheroid structures that retain stem cell characteristics. AECs did not require other cell-derived feeder layers to maintain Oct-4 expression, did not express telomerase and are non-tumorigenic upon transplantation. Based on immunohistochemical and genetic analysis, had the potential to differentiate to all three germ layers: endoderm (liver, pancreas), mesoderm (cardiomyocyte) and ectoderm (neural cells) in vitro (Miki et al., 2005). Sarkar et al. (1980) have also shown that HAECs obtained from human placenta were able to survive into the transection cavities in the spinal cord of bonnet monkeys, support the growth of host axons through them and prevent glial scar formation at the cut ends, and may prevent death in axotomized cells or attract the growth of new collateral sprouting. They speculated that HAECs may possess certain properties equal to the beneficial effects of neural tissue in repairing spinal cord injury. Apart from this speculation, there are two more reasons why HAEC transplantation studies are warranted to understand the long-term effects of such transplantations. First, there was no evidence of immunological rejection, probably due to the non-antigenic nature of the HAECs. Second, unlike neural tissue, procurement of HAEC does not involve legal or ethical problems (Sakuragawa et al., 1996; Elwan and Sakuragawa, 1997; Sakuragawa et al., 2000; Takahashi et al., 2002).

18.5

Amniotic fluid-derived stem cells

We have recently described a pluripotent population of cells derived from amniotic fluid defined as ‘amniotic fluid stem (AFS) cells’. Similar cells can be derived from chorionic villi (placenta) samples. The following paragraphs will describe in detail their isolation, characterization and differentiation in vitro into different lineages (De Coppi et al., 2007).

18.5.1 Cell preparation and culture methods Chorionic villi samples and human amniotic fluid was obtained following informed consent at 12–18 weeks of pregnancy from a total of 300 women

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between 23 and 42 years of age. In order to rule out the presence of maternal cells, only amniotic samples of male fetuses were used. Karyotypic analyses of the AFS cells showed an xy phenotype in all cells. Samples were seeded in a 22 mm × 22 mm cover slip in a volume of 2 ml and grown to confluence for 3–4 weeks at 95% humidity and 37 °C. Fresh medium was applied after 5 days of culture and every third day thereafter. The culture medium consisted of modified Eagle’s medium (GIBCO/BRL, Grand Island, NY), 18% Chang Medium B (Irvine Scientific, Santa Ana, CA), 2% Chang C (Irvine Scientific) with 15% embryonic stem cell-certified fetal bovine serum (ES-FBS; GIBCO/BRL), 1% antibiotics (GIBCO/BRL) and l-glutamine (Sigma-Aldrich, St Louis, MO). The cells were subcultured using 0.25% trypsin containing 1 mM EDTA for 5 minutes at 37 °C.

18.5.2 Isolation and characterization of amniotic fluid-derived stem cells The isolation of the AFS cells was performed using the surface antigen c-Kit (CD117), the receptor for stem cell factor (Zsebo et al., 1990). We used immunoselection with magnetic microspheres to isolate the c-Kit-positive cells; 1% of the amniotic cells were positive for c-Kit showing a doubling time of about 36 hours. We have routinely established clonal AFS cell lines as described in Section 18.5.3 (Takeda et al., 1992; Mosquera et al., 1999). These cells grew without the use of a feeder layer. These c-Kit-positive cell populations were termed amniotic fluid-derived stem (AFS) cells. We used flow cytometry to assess markers expressed by human AFS cells. Five clonal lines gave similar results. The cells were positive for Class I MHC antigens (HLA-ABC), and some were weakly positive for MHC Class II (HLA-DR). The AFS cells were negative for markers of the hematopoietic lineage (CD45) and of hematopoietic stem cells (CD34, CD133). However, they stained positively for a number of surface markers characteristic of mesenchymal and/or neural stem cells, but not embryonic stem (ES) cells, including CD29, CD44 (hyaluronan receptor), CD73, CD90 and CD105 (endoglin) (Barry et al., 1999; Barry et al., 2001; Tsai et al., 2004). Human AFS cells also were positive for stage-specific embryonic antigen (SSEA)-4 (Kannagi et al., 1983), a marker expressed by ES cells but generally not by adult stem cells. The AFS cells did not express other surface markers characteristic of ES (Thomson et al., 1998; Carpenter et al., 2003) and embryonic germ (EG) cells (Shamblott et al., 1998), SSEA-3 and Tra-1-81. Some lines were weakly positive for Tra-1-60. Over 90% of the cells expressed Oct-4, a transcription factor, which has been associated with the maintenance of the undifferentiated state and the pluripotency of ES and EG cells (Pan et al., 2002). Whereas normal somatic stem cells are non-tumorigenic, ES cells grow as teratocarcinomas when implanted in vivo (Evans and Kaufman, 1981;

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Martin 1981; Cowan et al., 1998; Thomson et al., 1998). None of 4 human AFS cell lines tested, including late-passage cells, formed tumors in severe combined immunodeficient (SCID)/Beige mice (CB17/lcr.Cg-PrkdcscidLystbg/Crl). Moreover, karyotypes of 11 human lines from pregnancies in which the fetus was male revealed one X and one Y chromosome and a normal diploid complement of autosomes. There were no major chromosomal rearrangements as judged by Giemsa banding. The AFS cells, even after expansion to 250 population doublings (pd), showed a homogeneous, diploid DNA content in the G1 phase of the cell cycle. The G1 and G2 cellcycle checkpoints appeared intact. Analysis of terminal restriction fragments (Bryan et al., 1998) showed that the average length of telomeres of human AFS cells stayed constant, B20 kbp, between early (20 pd) and late passage (250 pd) cells.

18.5.3 Cell cloning and retroviral marking In order to investigate whether the placenta and amniotic fluid contain stem cells that are able to differentiate into multiple lineages, cell colonies derived from single cells were expanded. The AFS cells were subsequently infected with CMMP-eGFP40 and cell clones that expressed green fluorescent protein (GFP) were obtained by limiting dilution. Clonality of the AFS cells was confirmed by Southern blot analysis with a GFP probe. Genomic DNA was extracted from 1 × 107 cells by phenol-chloroform extraction. After digesting 10 mg of DNA overnight at 37 °C with BamHI, an electrophoresis in a 1% agarose gel was performed. The separated fragments were denatured with alkali, transferred on to membranes (Zeta-Probe GT, BioRad) and incubated with a 32P-labeled GFP DNA probe overnight at 65 °C in Rapid-hyb buffer (Amersham Biosciences). Finally, the hybridized blots were washed according to the manufacturer’s instructions. Bound probe was visualized by autoradiography.

18.5.4 Multilineage induction Clonal lineages from different patients underwent adipogenic, osteogenic, myogenic, endothelial, neurogenic and hepatogenic differentiation in vitro. Adipogenic differentiation AFS cells were seeded at a density of 3000 cells/cm2 and were cultured in Dulbecco’s modified Eagle’s medium (DMEM) low-glucose medium with 10% FBS, antibiotics (Pen/Strep, Gibco/BRL) and adipogenic supplements (1 μM dexamethasone, 1 μM 3-isobutyl-1-methylxanthine, 10 μg/ml insulin, 60 μM indomethacin (Sigma-Aldrich). The cells changed their phenotype

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from elongated to round within 8 days. This coincided with the accumulation of intracellular triglyceride droplets. After 16 days in culture, more than 95% of the cells had their cytoplasm completely filled with lipid-rich vacuoles, which stained positively with Oil-O-Red. The c-kit-negative selected cells did not show any phenotypic change and did not stain with Oil-O-Red under the adipogenic conditions after 16 days of culture. Adipogenic differentiation was confirmed by RT-PCR analysis. The expression of both peroxisome proliferation-activated receptor γ 2 (pparγ2), a transcription factor that regulates adipogenesis, and lipoprotein lipase were up-regulated in the AFS cells under adipogenic conditions (Kim et al., 1998; Rangwala et al., 2000). AFS cells under control conditions and negatively selected cells in adipogenic medium did not express either gene at any time point. AFS cells were seeded on polyglycolic acid (PGA) polymer scaffolds at a density of 10 × 106 cells/cm2. Cells were induced into an adipogenic lineage in a bioreactor for 16 days. The scaffolds were implanted subcutaneously in athymic mice, harvested after 4 and 8 weeks, and analyzed. The retrieved scaffolds showed the formation of fatty tissues grossly. The presence of adipose tissue was confirmed with Oil-O-Red staining. Osteogenic differentiation Cells were seeded at a density of 3000 cells/cm2 and were cultured in DMEM low-glucose medium with 10% FBS (Gibco/BRL), Pen/Strep and osteogenic supplements (100 nM dexamethasone, 10 mM betaglycerophosphate (Sigma-Aldrich), 0.05 mM ascorbic acid-2-phosphate (Wako Chemicals)). Light microscopy analysis showed that the AFS cells developed an osteoblastic-like appearance with finger-like excavations into the cell cytoplasm within 4 days in osteogenic medium (Karsenty, 2000; Olmsted-Davis et al., 2003). After 16 days the cells aggregated in the typical lamellar bonelike structures and increased their expression of alkaline phosphatase (AP). Calcium accumulation was evident after 1 week and increased over time. In order to confirm the cytochemical findings, AP activity was measured using a quantitative assay, which measured p-nitrophenol, equivalent to AP production. The AFS cells showed more than a 200 times increase in AP production in the osteogenic-inducing medium compared with cells grown in control medium at days 16 and 24. After that time the levels of AP decreased. No AP production was detected in c-kit-negative selected amniotic cells cultured in osteogenic medium at any time point. AP expression was confirmed at the RNA level. No activation of the AP gene was detected at 8, 16, 24 and 32 days in the AFS cells grown in control medium. In contrast, AFS cells grown in osteogenic medium showed an activation of the AP gene at each time point. Expression of Runx2, a transcription factor for osteogenic induction (Lian and Stein, 2003), was highly induced in cells

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grown in osteogenic medium at day 8. The expression of Runx2 in the controls was significantly lower. Osteocalcin was also expressed in the AFS cells in osteogenic conditions at day 8. No expression of ostecalcin was detectable in the AFS cells in the control medium and in the negatively selected cells in the osteogenic conditions. A major feature of osteogenic differentiation is the ability of the cells to precipitate calcium. Cell-associated mineralization can be analyzed using von Kossa staining and by measuring the calcium content of cells in culture. Von Kossa staining of the AFS cells grown in the osteogenic medium showed enhanced silver nitrate precipitations by day 16, indicating high levels of calcium. Calcium precipitation continued to increase exponentially at 24 and 32 days. In contrast, cells in the control medium did not form silver nitrate precipitations after 32 days. Microscopic examination of stained cells showed no calcification in the osteogenic medium treated cells at days 4 or 8, but strong black silver nitrate precipitates were noticed in osteogenic induced cells after 16, 24 and 32 days in culture. In the AFS cells cultured in control medium, no precipitates were noticed over the 32 day time period. Calcium deposition by the cells was also measured with a quantitative chemical assay, which measures calcium–cresolophthalein complexes. AFS cells undergoing osteogenic induction showed a significant increase in calcium precipitation after 16 days (up to 4 mg/dl). The precipitation of calcium increased up to 70 mg/dl at 32 days. In contrast, AFS cells grown in control medium did not show any increase in calcium precipitation (1.6 mg/dl) by day 32. In order to investigate the utility of AFS cells for tissue engineering, we embedded human AFS cells in an alginate/collagen scaffold by thermal inkjet printing (Roth et al., 2004; Xu et al., 2005). For the in vivo assessment, printed human AFS cell/scaffold constructs were incubated for 1 week in osteogenic medium and implanted subcutaneously in immunodeficient mice. An unseeded control scaffold was also implanted in each mouse. After 8 weeks, constructs were recovered and analyzed histologically using von Kossa’s stain. Highly mineralized tissue was observed from the implanted cell-seeded scaffolds but not from the implanted unseeded scaffolds. After 18 weeks the generation of hard tissue within the printed constructs was demonstrated by micro-computed tomography scanning of the recipient mice.The scaffolds containing AFS cells revealed blocks of bone-like material with density similar to the murine femoral bone. Control scaffolds lacking AFS cells did not promote the formation of bony tissue. Myogenic differentiation Cells were seeded at a density of 3000 cells/cm2 on plastic plates pre-coated with Matrigel (Collaborative Biomedical Products; incubation for 1 hour at

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37 °C at 1 mg/ml in DMEM) in DMEM low-glucose formulation containing 10% horse serum (Gibco/BRL), 0.5% chick embryo extract (Gibco/BRL) and Pen/Strep (Rosenblatt et al., 1995). At 12 hours after seeding, 3 μM 5-aza2-deoxycytidine (5-azaC; Sigma-Aldrich) was added to the culture medium for 24 hours. Incubation then continued in complete medium lacking 5-azaC, with medium changes every 3 days. Induction with 5-azacytidine for 24 hours promoted the formation of multinucleated cells over a 24–48 hour period. After 16 days, the AFS cells grown with myogenic medium formed myofiber-like structures that stained immunocytochemically with desmin and sarcomeric tropomyosin. AFS cells grown in control medium and cells negatively selected for c-kit cultured in myogenic medium did not demonstrate cell fusion or multinucleated cells. Only a few desmin-positive cells were present in the negatively selected amniotic cells with myogenic medium after 16 days. Expression of MyoD, Myf5, Myf6 (MRF4) and desmin, were analyzed using RT-PCR. MyoD and MRF4 were expressed by the AFS cells in culture at 8 days, and suppressed at 16 days. Both these genes were not expressed at 8 or 16 days in the controls. Desmin expression was induced at day 8 and increased by 16 days in the AFS cells cultured in myogenic medium. In contrast, there was no activation of desmin in the control cells at 8 and 16 days. Myf5 was present at 8 days and increased at 16 days in the AFS cells. Lower levels of the Myf5 gene were detected in the cells maintained in culture with the control medium at 16 days. The AFS cells were labeled with a fluorescence marker (PKH26 green fluorescent cell linker, Sigma-Aldrich) and were induced into a myogenic lineage. The myogenic cells were resuspended in rat tail collagen containing 17% Matrigel (BD Biosciences), were injected into the hindlimb musculature of athymic mice and were retrieved after 4 weeks. The injected myogenic cells showed the formation of muscle tissue which express desmin and maintained its fluorescence. Endothelial differentiation AFS cells were seeded at 3000 cells/cm2 on plastic plates pre-coated with gelatin and maintained in culture for 1 month in endothelial cell medium-2 (EG-MTM-2, Clonetics; Cambrex Bioproducts) supplemented with 10% FBS, and Pen/Strep. Recombinant human bFGF (StemCell Technologies) was added at intervals of 2 days at 2 ng/ml. After 1 week in culture the AFS cells started to change their morphology, and by the second week were mostly tubular. The cells stained positively for FVIII, KDR and P1H12. Moreover they expressed CD31 and VCAM once cultured in endothelial conditions. c-kit-negative selected cells cultured in the same conditions and AFS cells cultured in Chang medium for

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the same period of time were not able to form tubular structures and did not stain for endothelial-specific markers. The cells, once differentiated, were able to grow in culture for more than 1 month. Neurogenic differentiation AFS cells were seeded at a concentration of 3000 cells/cm2 on tissue culture plastic plates and cultured in DMEM low-glucose medium, Pen/Strep, supplemented with 2% dimethylsulfoxide (DMSO), 200 μM butylated hydroxyanisole (BHA; Sigma-Aldrich) and nerve growth factor (NGF; 25 ng/ml). After 2 days the cells were returned to AFS growth medium lacking DMSO and BHA but still containing NGF. Fresh NGF was added every 48 hours. For a two-stage induction procedure designed to yield dopaminergic cells, the AFS cells were first seeded on plates coated with fibronectin (1 μg/ml) and incubated in DMEM/F12 medium supplemented with N2 and 10 ng/ml bFGF for 8 days. Fresh bFGF was added every second day. Under these conditions over 80% of cells showed expression of nestin. The cells were then transferred to conditions that favor the production of dopaminergic neurons (Perrier et al., 2004). In a first differentiation phase, 80% of the initial AFS cell population became positive for nestin, a marker first defined in neural stem cells (Lehndahl et al., 1990). Under conditions expected to bias for dopaminergic neurons (Perrier et al., 2004), a fraction of these cells revealed a distinct pyramidal morphology. Transcript analysis showed expression of the GIRK2 gene, encoding a member of the G-protein-gated inwardly rectifying potassium (GIRK) channel family, a marker of dopaminergic neurons (Liao et al., 1996). Voltage clamping of individual cells revealed a bariumsensitive potassium channel consistent with the GIRK2 channel. Furthermore, after application of another neurogenic induction protocol using NGF,AFS cells acquired the ability to secrete the excitatory neurotransmitter l-glutamate in response to stimulation by potassium ions (Fig. 18.1). The ability to contribute to the development of the central nervous system was investigated by injecting human AFS cells (after induction in neurogenic medium) into the lateral cerebral ventricles of the brains of newborn mice, both wild-type and homozygous twitcher (C57BL/6J-Galctwi) mutants (Taylor et al., 2006). The twitcher mice are deficient in the lysosomal enzyme galactocerebrosidase and undergo extensive neurodegeneration and neurological deterioration, starting with dysfunction of oligodendrocytes, similar to that seen in the genetic disease Krabbe globoid leukodystrophy (Suzuki and Suzuki, 1995). Implantation into the lateral ventricles affords the cells access to the subventricular zone, a secondary germinal zone in the cerebrum that persists throughout life and from which resident neural progenitors readily migrate into and integrate within cerebral parenchyma (Taylor

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Neuronal

10

Urothelium

Heat inactivated induced 6 days

Induced 6 days

Not induced 6 days

Induced 4 days

Not induced 4 days

Induced 2 days

0 Not induced 2 days

Glutamic acid secretion (ng/cell hour)

20

18.1 Only the progenitor cells cultured under neurogenic conditions showed the secretion of glutamic acid in the collected medium. The secretion of glutamic acid could be induced (20 minutes in 50 mM KCl buffer).

and Snyder, 1997). The grafted cells dispersed throughout the host mouse brains and survived for at least 2 months. After 1 month the human cells were present in different brain regions, including periventricular areas, the hippocampus and the olfactory bulb, where they integrated seamlessly and appeared morphologically indistinguishable from surrounding murine cells. Neither deformation of the host brain nor any neoplastic process was evident. The pattern of migration of the transplanted cells was not random; they populated the same areas in all engrafted mice. However, the number of engrafted human cells was higher in the brains of the twitcher mutants (70% of injected cells) than in the wild-type recipients (30%). Hepatic differentiation Cells were seeded at a density of 5000 cells/cm2 on plastic plates coated with Matrigel. They were expanded in AFS growing medium for 3 days to achieve a semi-confluent density. The medium was then changed to DMEM low-glucose formulation containing 15% FBS, 300 mM monothioglycerol (Sigma-Aldrich), 20 ng/ml hepatocyte growth factor (HGF, Sigma-Aldrich), 10 ng/ml oncostatin M (Sigma-Aldrich), 107 M dexamethasone (SigmaAldrich), 100 ng/ml FGF4 (Peprotech), 1 × ITS (insulin, transferrin, selenium; Roche) and Pen/Strep. The cells were maintained in this differentiation

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medium for 2 weeks, with medium changes every third day. They were then harvested using trypsin, and plated into a collagen sandwich gel (0.11 mg/ cm2 for both the lower and upper layers) (Woodbury et al., 2000; Schwartz et al., 2002). The hepatogenic cultures were maintained for up to 45 days. We observed induction of markers characteristic of the hepatocyte lineage (endoderm) in AFS cells cultured for up to 45 days in a multi-step system. Proteins expressed by hepatic differentiated AFS cells included albumin, alpha-fetoprotein, hepatocyte nuclear factor 4 (HNF4), hepatocyte growth factor receptor (c-Met) and the multidrug resistance membrane transporter MDR1. Furthermore, hepatic lineage cells obtained by differentiation of AFS cells were able to secrete urea, a characteristic liverspecific function.

18.6

Conclusions

Fetal tissue has been used in the past for transplantation and tissue engineering research because of its pluripotency and proliferative ability. Fetal cells maintain a higher capacity to proliferate than adult cells and may preserve their pluripotency longer in culture. However, fetal cell transplants are plagued by problems that are very difficult to overcome. Beyond the ethical concerns regarding the use of cells from aborted fetuses or living fetuses, there are other issues that remain a challenge. Previous studies have shown that it takes almost six fetuses to provide enough material to treat one patient with Parkinson’s disease. In the last few yeas it has been shown that placenta and amniotic fluid contain a large variety of cells. The majority of cells in the placenta and in the amniotic fluid are already differentiated. Therefore these cells possess a limited proliferative ability. In this chapter we have summarized the different progenitors that have been described in the last few years to be present in both amniotic fluid and placenta. In particular we have focused our attention on MSCs and AFS cells, isolated using c-kit cells, and their different proliferative and differentiative potential. In conclusion, placenta and amniotic fluid could be an excellent cell source for therapeutic applications. When compared with ES cells, fetal stem cells are easily differentiated into specific cell lineages, they do not need any feeder layer to grow, and avoid the current controversies associated with the use of human ES cells.

18.7

References

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barry fp, boynton re, haynesworth s, murphy jm, zaia j. The monoclonal antibody SH-2, raised against human mesenchymal stem cells, recognizes an epitope on endoglin (CD105). Biochem Biophys Res Commun. 1999;265:134–9. barry f, boynton r, murphy m, haynesworth s, zaia j. The SH-3 and SH-4 antibodies recognize distinct epitopes on CD73 from human mesenchymal stem cells. Biochem Biophys Res Commun. 2001;289:519–24. brace ra, resnik r. Dynamics and disorders of amniotic fluid. In Creasy RK & Renik R (eds), Maternal–Fetal Medicine. Philadelphia: Saunders, 1999, pp. 623–43. bryan tm, englezou a, dunham ma, reddel, rr. Telomere length dynamics in telomerase-positive immortal human cell populations. Exp Cell Res. 1998;239:370–8. carpenter mk, rosler e, rao, ms. Characterization and differentiation of human embryonic stem cells. Cloning Stem Cells. 2003;5:79–88. cousineau j, potier m, dallaire l, melancon sb. Separation of amniotic fluid cell types in primary culture by Percoll density gradient centrifugation. Prenat Diagn. 1982 Oct;2(4):241–9. cowan ca, klimanskaya i, mcmahon j, atienza j, witmyer j, zucker jp, wang s, morton cc, mcmahon ap, powers d, melton da. Derivation of embryonic stem-cell lines from human blastocysts. N Engl J Med. 2004;350:1353–6. crane jp, cheung sw. An embryogenic model to explain cytogenetic inconsistencies observed in chorionic villus versus fetal tissue. Prenat Diagn. 1988;8:119–29. de coppi p, bartsch g jr, siddiqui mm, xu t, santos cc, perin l, mostoslavsky g, serre ac, snyder ey, yoo jj, furth m, soker s, atala a. Isolation of clonal amniotic stem cell lines with potential for therapy. Nat Biotechnol. 2007 Jan;25(1):100–6. elwan ma, sakuragawa n. Evidence for synthesis and release of catecholamines by human amniotic epithelial cells. Neuroreport. 1997;8:3435–8. evans mj, kaufman mh. Establishment in culture of pluripotential cells from mouse embryos. Nature 1981;292:154–6. fuchs jr, kaviani a, oh jt, lavan d, udagawa t, jennings rw, wilson jm, fauza do. Diaphragmatic reconstruction with autologous tendon engineered from mesenchymal amniocytes. J Pediatr Surg. 2004 Jun;39(6):834–8; discussion 834–8. gosden cm. Amniotic fluid cell types and culture. Brit Med Bull. 1983;39:348–54. haigh t, chen c, jones cj, aplin jd. Studies of mesenchymal cells from 1st trimester human placenta: expression of cytokeratin outside the trophoblast lineage. Placenta. 1999;20:615–25. hoehn h, salk d. Morphological and biochemical heterogeneity of amniotic fluid cells in culture. Methods Cell Biol. 1982;26:11–34. in‘t anker ps, scherjon sa, kleijburg-van der keur c, noort wa, claas fh, willemze r, fibbe we, kanhai hh. Amniotic fluid as a novel source of mesenchymal stem cells for therapeutic transplantation. Blood. 2003 Aug 15;102(4):1548–9. in‘t anker ps, scherjon sa, kleijburg-van der keur c, de groot-swings gm, claas fh, fibbe we, kanhai hh. Isolation of mesenchymal stem cells of fetal or maternal origin from human placenta. Stem Cells. 2004;22(7):1338–45. kannagi r, cochran na, ishigami f, hakomori s, andrews pw, knowles bb, solter d. Stage-specific embryonic antigens (SSEA-3 and -4) are epitopes of a unique globo-series ganglioside isolated from human teratocarcinoma cells. EMBO J. 1983;2:2355–61. karsenty, g. Role of Cbfa1 in osteoblast differentiation and function. Semin Cell Dev Biol. 2000;11:343–6.

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kaviani a, perry te, dzakovic a, jennings rw, ziegler mm, fauza do. The amniotic fluid as a source of cells for fetal tissue engineering. J Pediatr Surg. 2001;36:1662–5. kaviani a, guleserian k, perry te, jennings rw, ziegler mm, fauza do. Fetal tissue engineering from amniotic fluid. J Am Coll Surg. 2003;196:592–7. kim jb, wright hm, wright m, spiegelman bm. ADD1/SREBP1 activates PPARgamma through the production of endogenous ligand. Proc Natl Acad Sci USA. 1998;95:4333–7. lendahl u, zimmerman lb, mckay, rd. CNS stem cells express a new class of intermediate filament protein. Cell. 1990;60:585–95. li cd, zhang wy, li hl, jiang xx, zhang y, tang p, mao n. Isolation and identification of a multilineage potential mesenchymal cell from human placenta. Placenta. 2005a Sep 17. [Epub ahead of print]. li cd, zhang wy, li hl, jiang xx, zhang y, tang ph, mao n. Mesenchymal stem cells derived from human placenta suppress allogeneic umbilical cord blood lymphocyte proliferation. Cell Res. 2005b Jul;15(7):539–47. lian jb, stein gs. Runx2/Cbfa1: a multifunctional regulator of bone formation. Curr Pharm Des. 2003;9(32):2677–85. liao yj, jan yn, jan ly. Heteromultimerization of G-protein-gated inwardly rectifying K+ channel proteins GIRK1 and GIRK2 and their altered expression in weaver brain. J. Neurosci. 1996;16:7137–50. lotgering fk, wallenburg hc. Mechanisms of production and clearance of amniotic fluid. Semin Perinatol. 1986 Apr;10(2):94–102. martin gr. Isolation of a pluripotent cell line from early mouse embryos cultured in medium conditioned by teratocarcinoma stem cells. Proc Natl Acad Sci USA. 1981;78:7634–8. medina-gomez p, johnston th. Cell morphology in long-term cultures of normal and abnormal amniotic fluids. Hum Genet. 1982;60(4):310–13. miki t, lehmann t, cai h, stolz db, strom sc. Stem cell characteristics of amniotic epithelial cells. Stem Cells. 2005 Nov–Dec;23(10):1549–59. milunsky a. Amniotic fluid cell culture. In Milunsky A (ed.), Genetic Disorders and the Fetus. New York: Plenum Press, 1979, pp. 75–84. mosquera a, fernández jl, campos a, goyanes vj, ramiro-di´az j, gosálvez j. Simultaneous decrease of telomerase length and telomerase activity with ageing of human amniotic fluid cells. J Med Genet 1999;36:494–6. okawa h, okuda o, arai h, sakuragawa n, sato k. Amniotic epithelial cells transform into neuron-like cells in the ischemic brain. Neuroreport. 2001 Dec 21;12(18):4003–7. olmsted-davis ea, gugala z, camargo f, gannon fh, jackson k, kienstra ka, shine hd, lindsey rw, hirschi kk, goodell ma, brenner mk, davis ar. Primitive adult hematopoietic stem cells can function as osteoblast precursors. Proc Natl Acad Sci USA. 2003;100:15877–82. pan gj, chang zy, schöler hr, pei d. Stem cell pluripotency and transcription factor Oct4. Cell Res. 2002;12:321–9. perrier al, tabar v, barberi t, rubio me, bruses j, topf n, harrison nl, studer l. Derivation of midbrain dopamine neurons from human embryonic stem cells. Proc Natl Acad Sci USA. 2004;101:12543–8. rosenblatt jd, lunt ai, parry dj, partridge ta. Culturing satellite cells from living single muscle fiber explants. In Vitro Cell Dev Biol Anim. 1995;31:773–9.

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19 The use of adipose progenitor cells in urology D. S. DAV É and L. V. R O D R Í G U E Z, University of California Los Angeles, USA

Abstract: Mesenchymal stem cells within human adipose tissue are easily procured in abundant quantities, exhibit multilineage differentiation, and can be readily transplanted into an autologous host. Referred to as adipose-derived stem cells, these cells have a characteristic molecular profile and phenotype. Lacking the ethical and political limitations associated with embryonic stem cells, adipose-derived stem cells represent a promising stem cell source for the field of tissue engineering and regenerative medicine. This chapter reviews the current state-of-theart knowledge with respect to the use of adipose-derived stem cells for tissue engineering with a particular focus on applications within the field of urology. Key words: adipose tissue, adult stem cells, tissue engineering, urinary tract reconstruction.

19.1

Introduction

The goal of tissue engineering as a field is the regeneration and restoration of damaged tissues to a normal functional state by incorporating cell-based therapies and using principles of material science. Current strategies may employ acellular matrices which direct the body’s ability to regenerate by allowing for the ingrowth of surrounding cells. It is unclear as to whether this regeneration results from proliferation of differentiated smooth muscle cells (SMCs) or by migration and proliferation of smooth muscle (SM) progenitor cells, which are lineage-restricted cells designed to replenish damaged SMCs in the case of injury. In addition, cell-based therapies with or without the use of matrices may be employed, in which case regeneration results from direct replication of introduced cells, which may either be reimplanted directly after harvest or after in vitro expansion in specific culture conditions. Although cells utilized in this manner may be heterologous, allogenic, or autologous, the ideal source would involve the use of autologous cells as this avoids the possibility of rejection and obviates the need for immunosuppressive therapy. For urological applications, autologous cells have been obtained from SM and urothelial biopsies of the 395

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urinary tract. The already differentiated cells within the biopsy specimen have limited proliferative capability due to cell senescence. Although the SM progenitor cells within the specimen have improved longevity, the fact that they may be derived from the diseased organ of the host makes them a less-than-ideal source of cells for tissue engineering applications. In addition, limited cell yield may pose further restrictions. The ideal cells for tissue engineering should be easily procured using minimally invasive techniques, provide abundant quantities of cells derived from normal, healthy tissue, proliferate quickly in a well-regulated manner, exhibit multilineage differentiating capabilities for regeneration of multiple tissue types, and be capable of being readily transplanted into an autologous host. Stem cells represent such a self-renewing population of unspecialized cells that can differentiate into multiple cell types. Adult stem cells that can differentiate into mesodermal tissues, also called mesenchymal stem cells (MSCs), have been identified within several organs – including fetal liver, umbilical cord blood, and bone marrow stroma (Kern et al., 2006, Wagner et al., 2005). These cells can be stimulated to differentiate into adipocytes, chondrocytes, myoblasts, and osteoblasts. Although they lack the political and ethical limitations of embryonic stem cells, disadvantages of MSCs include morbidity of cell procurement and a relatively low yield of stem cells. Recently, multipotent MSCs have been identified within the stromal vascular fraction of human adipose tissue. Clonal studies have shown that these cells exhibit multilineage differentiation into adipogenic, chondrogenic, myogenic, osteogenic, and neurogenic cells, leading many to refer to these cells as adipose-derived stem cells (ASCs) (Zuk et al., 2001, Zuk et al., 2002). These cells do not exhibit the ethical or political limitations seen with embryonic stem cells. In addition, they are generally found in abundance and are easily harvested with little morbidity, making them promising candidates for use in tissue engineering. Furthermore, their ability to differentiate into SMCs as well as vascular and neurological structures makes them an ideal candidate for use in urology. This chapter will review the current knowledge regarding the state-of-the art use of adipose progenitor cells in the field of tissue engineering with a specific focus on SM differentiation for the purpose of urological applications. In particular, we will discuss: • • • • • •

nomenclature and origin of adipose progenitor cells; isolation procedures; molecular characterization; differentiation capacity of adipose-derived stem cells; applications in the field of urology; future directions and roles for adipose progenitor cells.

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19.2

397

Nomenclature and origin of adipose progenitor cells

There is some confusion in the literature regarding the nomenclature that should be used to describe the plastic-adherent, multipotent precursor cells derived from adipose tissue. Adipose tissue originates within the embryonic mesoderm, with earliest evidence of adipocyte formation sometime during the second trimester of development (Martin et al., 1998, Nnodim, 1987). The actual origin of progenitor cells found within adipose tissue, however, remains controversial, as evidenced by the myriad of names in the literature used to describe these cells: processed lipoaspirate (PLA) cells, adipose tissue-derived stromal cells, preadipocytes, pericytes, lipoblasts, and adipose stroma vascular cell fraction to name a few. It has not yet been proven whether these cells originate within the endothelial, pericyte, or stromal compartments (Wright and Hausman, 1990a, Wright and Hausman, 1990b). With the presence of similar cell surface antigens, these cells seem to share a common origin with endothelial cells (Wright and Hausman, 1990a, Wright and Hausman, 1990b). Alternatively, recent studies using genetically engineered murine models to track transplanted bone marrow-derived cells have shown that cells similar to colony-forming-unit fibroblasts (CFU-F) can integrate into extramedullary fat depots where they can differentiate into adipocytes under the influence of thiazolidinediones or a high fat diet (Crossno et al., 2006). This raises the possibility that adipose-derived progenitor cells represent a circulating pool of progenitor cells which originate in the bone marrow. Further confusion regarding the appropriate nomenclature for adiposederived progenitor cells comes from the method by which these cells are isolated, namely by collagenase digestion of raw lipoaspirate. Because the progenitor cells obtained in this fashion represent only a fraction of the total cultured population of cells, many have proposed the use of the term PLA to describe these cells. Regardless of the exact origin of these cells or the exact percentage of total cultured cells these cells represent, it is clear that progenitor cells within the adipose stromal vascular fraction exhibit multipotent differentiation. Although the nomenclature of these cells remains a debatable topic, it has been determined by an International Fat Applied Technology Society consensus meeting to refer to these cells as ‘adipose-derived stem cells’ (ASCs). For the purpose of this chapter, therefore, we will henceforth refer to these cells as adipose-derived stem cells (ASCs).

19.3

Isolation procedures

The initial methods outlining the procurement of cells from adipose tissue were described in the 1960s (Rodbell, 1966a, Rodbell, 1966b, Rodbell and

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Jones, 1966). Since that time, we have learned that numerous factors affect the yield of viable cells, the proliferation rates of cells as well as the differentiation capacity of ASCs. Some of these factors include donor age, adipose tissue type, anatomic localization of tissue, surgical method of procurement, culturing environment, cell attachment surfaces, plating density, and media formulations. Adipose tissue has been typically classified as one of five types: bone marrow, brown, white, mammary, and mechanical; with each type serving a unique biological role. At the molecular level, it is well established that different fat depots exhibit distinct gene expression profiles resulting in differences in production of adipose-specific cytokines such as adiponectin, leptin, and resistin (Prunet-Marcassus et al., 2006, Rodbell, 1966a, Rodbell, 1966b, Rodbell and Jones, 1966, Schipper et al., 2006). These differences ultimately result in macroscopic differences in metabolic function – including fatty acid composition and storage, lipolytic activity – and perhaps differences in ASC numbers and differentiation capacity. Although data are limited, cellular composition and differentiation capacity of stem cells obtained from different fat depots have been shown to be notably different in a murine system, with greater cell yield and differentiation capacity seen within white adipose tissue compared with brown (Prunet-Marcassus et al., 2006). In humans, anatomic localization of adipose tissue seems to affect the yield of stem cells, with subcutaneous white adipose tissue from the arm contributing more cells compared with tissue obtained from the thigh, abdomen, and breast (Schipper et al., 2006). Stem cell yield also differs between subcutaneous and omental white adipose depots in humans (Van Harmelen et al., 2004). The unique characteristics of cells from different fat depots and their potential impact on tissue engineering applications deserve further scientific attention. Surgical methods used to procure ASCs have evolved over the years. With the advent of liposuction, the procurement process has been greatly simplified. The procedure of tumescent liposuction involves the subcutaneous instillation of a saline solution containing an anesthetic and/or epinephrine through a cannula. The lipoaspirate, which is a combination of the instilled solution and adipose tissue, is then suctioned out, with the size of removed, minced tissue particles determined by the size of the cannula used (Illouz, 1983). Using this method, viable ASCs have been identified within the lipoaspirate fluid as well as within the removed tissue with no apparent adverse effect on the number of viable cells due to the aspiration process (Lalikos et al., 1997, Moore et al., 1995, Oedayrajsingh-Varma et al., 2006). When compared with ultrasonic liposuction, adipose tissue harvested by resection or tumescent liposuction has a higher frequency of proliferating ASCs with a faster doubling time (Wagner et al., 2005). Although the age of the donor has an impact on attachment and proliferation capacity, with

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both factors significantly better in cells derived from younger donors, the ability to differentiate is maintained with ageing (Shi et al., 2005). Once isolated, ASCs are readily preserved under cryo conditions and expanded in culture with morphology similar to that of fibroblasts. Numerous factors – including plating density, culture media, time in passage, and plastic adherence – have been shown to affect the differentiation capacity of ASCs in vitro (Dicker et al., 2005, Lee et al., 2004). Differential expression of hundreds of genes was seen when ASCs were cultured under different media conditions (Lee et al., 2004). Low plating densities with the use of Dulbecco’s modified Eagle’s medium/MCDB media seem to facilitate adipogenic differentiation of ASCs (Dicker et al., 2005). In addition, supplementation of culture media with antioxidants and the use of low calcium concentrations have been shown to increase the growth rate and life span of ASCs in culture (Nakagami et al., 2006).

19.4

Molecular characterization

Although MSCs were initially isolated from bone marrow, it is now known that similar cell populations may be isolated from umbilical cord blood, peripheral blood, connective tissues of the dermis, skeletal muscle, and adipose tissue (Bieback et al., 2004, Erices et al., 2000, Goodwin et al., 2001, Jiang et al., 2002, Kern et al., 2006, Kuznetsov et al., 2001, Wagner et al., 2005, Zvaifler et al., 2000). Bone marrow-derived MSCs have been extensively studied and characterized at the molecular level. With this clear understanding and definition of the bone marrow MSC phenotype, significant advances have been made in the use of bone marrow stem cell transplantation, taking this treatment from a highly experimental procedure to standard therapy for several malignant and hereditary diseases (Ho and Punzel, 2003). In contrast, the phenotypic characterization of ASCs is still being established. It is clear, however, that a precise characterization of these cells will be paramount in establishing critical standards for future development and use of ASCs in clinical human applications. According to the Mesenchymal and Tissue Stem Cell Committee of the International Society for Cellular Therapy, a minimal set of four criteria must be met in order for a particular type of cell to be classified as an MSC: (a) they must be plastic adherent when maintained in standard culture conditions; (b) they must exhibit potential for osteogenic, adipogenic, and chondrogenic differentiation; (c) they must express CD73, CD90, and CD105; and finally (d) they must lack expression of hematopoietic lineage markers c-kit, CD14, CD11b, CD34, CD45, CD19, CD79α, and HLA-DR (Dominici et al., 2006). The surface immunophenotype of ASCs has been studied by several groups (Katz et al., 2005, Stashower et al., 1999, Williams et al., 1994, Yoshimura et al., 2006, Young et al., 1992, Zuk et al., 2001). ASCs

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Table 19.1 Molecular immunophenotype of human adipose-derived stem cells

Antigen type

Surface-positive antigen

Adhesion molecules

CD9 (tetraspan) CD29 (β1 integrin) CD49d (α4 integrin) CD54 (ICAM-1) CD105 (endoglin) CD166 (ALCAM) CD44 (Hyaluronate) CD71 (transferrin) CD10 (Common ALL antigen) CD13 (aminopeptidase) CD73 (5′ ecto-nucleotidase) Aldehyde dehydrogenase CD90 (Thy1) CD146 (Muc18) Collagen types I and III Osteopontin Osteonectin α-Smooth muscle actin Vimentin

Receptor molecules Enzymes

Extracellular matrix proteins

Cytoskeletal proteins Hematopoeitic markers Complement cascade Histocompatibility Antigen Stem cell markers Stromal cell markers

Surface-negative antigen CD11b (αb integrin) CD18 (β2 integrin) CD50 (ICAM-3) CD56 (NCAM) CD62 (E-selectin) CD104 (α4 integrin) CD16 (Fc receptor)

CD14, CD31, CD45 CD55 (decay accelerating factor) CD59 (protectin) HLA-ABC

HLA-DR

CD34 ABCG2 CD29, CD44, CD73, CD90, CD166

have been found to exhibit characteristic expression of a number of molecules that has been consistent between different laboratories despite differences in cell harvesting techniques or culture conditions. A list of these cell surface markers is presented in Table 19.1. When comparing the surface immunophenotypes of ASCs to bone marrow MSCs, there seems to be a high degree of similarity, with direct comparison revealing a greater than 90% identity between cells types (Zuk et al., 2001). In addition, the two cell types exhibit similar responses to treatment with tumor necrosis factor-α and beta-adrenergic agents (Dicker et al., 2005, Ryden et al., 2003).

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Molecular differences between the two cell types do occur, however. CD34, for example, which is required to be absent in order for the cell to be classified as an MSC, can be found on human ASCs in early passages (McIntosh et al., 2006, Pittenger et al., 1999). These differences have been used by a number of groups to enrich the ASC population obtained from the heterogeneous stromal vascular fraction by using either flow cytometry or immunomagnetic beads (Hutley et al., 2001, Miranville et al., 2004, Sengenes et al., 2005). In fact, one group contends that ASCs are contained within the CD34-containing cell population of the stromal vascular fraction (Miranville et al., 2004, Sengenes et al., 2005). At the DNA level, gene microarray technology has allowed for the comparison of gene expression profiles between ASCs and bone marrow MSCs. In a study analyzing the gene expression profiles of MSCs derived from adipose tissue, bone marrow, and umbilical cord blood and comparing these with fibroblasts, 25 genes were found to be upregulated in MSCs compared with fibroblasts, with no phenotypic differences seen between the three stem cell preparations when using a panel of 22 surface antigens (Wagner et al., 2005). In contrast, several hundred expressed sequence tags were found to be differentially expressed between ASCs and the other two stem cell preparations (Wagner et al., 2005). Another study identified 24 genes that were differentially upregulated in ASCs compared with bone marrow MSCs. This was in conjunction with 8 surface marker proteins that were found to be differentially expressed between the two cell types (Lee et al., 2004). This group demonstrated that less than 1% of genes are felt to be differentially expressed between ASCs and bone marrow MSCs. Yet another study showed no significant difference in the expression of 28 selected genes between the two cell types (Winter et al., 2003). In a more comprehensive study using gene chip technology, another group concluded that human ASCs and bone marrow MSCs share a common transcriptome (Fraser et al., 2006). Proteomic studies have also confirmed the similarities between ASCs and bone marrow MSCs. A direct comparison of the two cell types using two-dimensional gel electrophoresis has identified 19 distinct proteins with a greater than 1.5-fold difference in expression between the two (R. Izadpanah, B. Bunnell, C. Kreidt, I. Kheterpal, A. White, unpublished observation, 2006). These data seem to support the hypothesis that ASCs and bone marrow MSCs share a common precursor cell; however, conclusive evidence to corroborate this hypothesis is still lacking. Regardless, it is universally accepted that further detailed molecular characterization of ASCs will be necessary in order to provide common standards and a precise definition of what truly represents an ASC. This will be important in establishing protocols that may be standardized between laboratories in order to facilitate the future use of ASCs in clinical applications.

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19.5

Differentiation capacity of adiposederived stem cells

The majority of the urinary tract is lined by SM and uroepithelium that is highly vascularized and innervated. When injury occurs to these tissues, normal form and function becomes altered. Cell-based therapies are designed to restore the normal capacity of these tissues by reintroducing a population of undamaged cells into damaged tissues, thereby allowing for the regeneration of healthy, functional tissue. Strategies utilizing tissue biopsies from the affected organ for in vitro cell expansion followed by reintroduction of the expanded cell population are limited by the fact that these cells are derived from a diseased organ. For this reason, the use of stem cells that can be differentiated into the desired tissue type in an in vitro setting is alluring. For urological applications, the ability to differentiate stem cells into SMCs, vascular structures, and nerves would provide a great foundation for the reconstruction and restoration of the urinary tract. SMCs represent a heterogeneous population of cells that share common features such as a spindle shape, low proliferative rate, reduced synthesis of extracellular matrix, and ligand-induced contractility (Small and Gimona, 1998, Yang et al., 1999). These cells exist in a wide range of differentiated states ranging from immature cells to mature SMCs characterized by their full expression of the cytoskeletal and contractile proteins that give them their identity. It is the sequential expression of these protein markers that confers the diversity of phenotypes ultimately leading to their heterogeneity. In order to establish that a particular cell is a mature SMC, certain unique markers need to be present. These markers include alpha-smooth muscle actin (ASMA), SM-22, calponin, caldesmon, smoothelin, and SM myosin heavy chain (MHC). The expression of ASMA, which is an early marker of SM differentiation, does not in itself define an SM lineage. The other markers of SM differentiation, however, are more highly restricted to differentiated SM and, in particular, SM MHC and smoothelin are not detected in any cell type other than contractile SMCs (Miano, 2002, van der Loop et al., 1996). MSCs derived from bone marrow have been shown to exhibit the capacity for osteogenic, chondrogenic, myogenic, adipogenic, and neurogenic differentiation in the presence of lineage-specific induction factors. Several studies have shown a similar capacity for multilineage differentiation exhibited by ASCs (Erickson et al., 2002, Planat-Benard et al., 2004a, Safford et al., 2001, Zuk et al., 2001, Zuk et al., 2002). Several groups have specifically examined the potential for SM differentiation of MSCs (Kashiwakura et al., 2003, Orlic et al., 2001, Religa et al., 2002, Wolf et al., 2003). If bone marrowderived MSCs exhibit the potential for SM differentiation, it would follow that ASCs should harbor the same potential. Indeed, human ASCs can

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express ASMA, calponin, and SM-22 in vitro, consistent with an SM phenotype (Jeon et al., 2006, Lee et al., 2006). In addition, the expression of SM-associated markers can be enhanced by the addition of transforming growth factor-β and sphingosylphosphorylcholine (Jeon et al., 2006, Lee et al., 2006). Human ASCs cultured in SM induction media (SMIM) differentiate into an SM phenotype characterized by the genetic expression of the early and late markers of SM differentiation ASMA, SM-22, calponin, caldesmon, smoothelin, and SM MHC (Rodriguez et al., 2006). This was confirmed by noting increased protein expression of the same markers. In order to rule out the possibility that the SMIM selected out a subpopulation of progenitor cells, similar differentiation experiments were carried out on clonal populations of cells derived from the PLA. SM differentiation occurred in all clonal populations. In addition, it was shown that all clonal populations maintained the capacity for trilineage differentiation (osteogenic and adipogenic, in addition to leiomyogenic), corroborating the presence of a multilineage precursor cell within adipose tissue. The functional goal of SMCs is to contract, and this ability provides functional proof that myocytes are in a fully differentiated state. Using collagen gel contraction assays, we demonstrated that ASCs differentiated into SMCs exhibit dose-dependent contraction in response to the muscarinic agonist carbachol that could be blocked by the addition of atropine. Furthermore, this contractility was absent when evaluating precursor cells prior to differentiation. It is clearly established that SM differentiation is not only regulated by chemical factors, but also by mechanical factors such as mechanical strain, which has been shown to modulate SMC phenotype, function, and matrix remodeling (Li and Xu, 2000, Owens, 1995, Thyberg, 1996). Previous studies have shown that cyclic mechanical strain plays an important role in the remodeling of the vascular wall in vivo and in the remodeling of tissueengineered vascular grafts in vitro (Kim et al., 1999, Li and Xu, 2000, Niklason et al., 1999, Seliktar et al., 2000, Williams, 1998). The effect of mechanical stimulation on the SM differentiation of MSCs has not clearly been defined, and few studies have been undertaken to delineate these effects (Hamilton et al., 2004, Kurpinski et al., 2006, Park et al., 2004). Data evaluating the effects of mechanical strain on the SM differentiation of ASCs are even more limited. One study concluded that cyclic uniaxial strain inhibited proliferation, caused alignment of cells in a direction perpendicular to the axis of strain, and resulted in decreased expression of the early SMC markers ASMA and calponin (Lee et al., 2007). The type and direction of strain applied as well as the alignment of cells with respect to the axis of strain seem to play an important role in SM differentiation. Further studies specifically delineating these factors and their effects on the SM differentiation of ASCs are needed.

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As all urological tissues and organs are highly vascularized and innervated, differentiation of ASCs into vascular and neurological structures in addition to SMCs would provide an improved ability to regenerate functional urological tissues. Indeed, ASCs and bone marrow-derived MSCs have a similar potential for endothelial differentiation (Urbich and Dimmeler, 2004). ASCs are known to produce numerous pro-angiogenic cytokines and growth factors including hepatocyte growth factor (HGF), transforming growth factor-β (TGF-β), vascular endothelial growth factor (VEGF), placental growth factor, fibroblast growth factor-2 (FGF-2) and angiopoietin-1 (Cao et al., 2005, Nakagami et al., 2005, Nakagami et al., 2006, Rehman et al., 2004). In an in vivo setting, injection of ASCs in murine models promotes vessel incorporation, neovascularization following induction of ischemia, and formation of vessel-like structures (Moon et al., 2006, Planat-Benard et al., 2004b). Under special neuroinductive conditions, ASCs can differentiate into cells expressing neuronal differentiation markers such as type III β-tubulin, glial fibrillary acidic protein (GFAP), nestin, NeuN, and intermediate filament M (Romanov et al., 2005, Safford et al., 2002). When induced to undergo neural differentiation, predifferentiated ASCs injected into the brains of rats following artificially induced ischemic brain injury led to improved motor recovery and functional deficits (Kang et al., 2003). The ability to safely and reliably differentiate ASCs into SMCs as well as vascular and neurological structures in an in vitro setting has recently been clearly established. Realizing the theoretical advantages of using SMCs, vascular and neuronal cells derived from a population of healthy stem cells that can be easily procured in abundant quantities will lay the foundation for the use of ASCs for tissue engineering applications across numerous fields of medicine including urology.

19.6

Applications in the field of urology

The urinary tract may be subjected to numerous potential injuries following fetal development that compromise normal cellular architecture and function. The majority of the urinary tract is composed of a combination of SM and uroepithelium. Traditional strategies for reconstruction of these tissues, when damaged, have employed tissue flaps and grafts from various parts of the body, intestinal segments, or synthetic materials to replace deficient tissues or to bolster weaknesses in native tissues. These strategies are limited in that they do not provide a means to recapitulate the full form or function of the native tissues and, in many instances, are associated with their own, unique complications or drawbacks – such as the formation of stones, strictures, mucus as well as a predisposition to metabolic abnormalities, urinary tract infections, and even de novo malignancies. Numerous advances in the

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past decade have brought tissue engineering to the forefront of scientific and medical research. A variety of strategies have been employed in the field of urology for the purpose of urinary tract reconstruction and regeneration. Although the use of tissue engineering in the field of urology is still in its infancy, its potential is vast and advances are being made at an exponential rate. The use of ASCs for the purpose of urinary tract reconstruction and SM regeneration offers numerous potential advantages over other current strategies that employ either SM biopsies from potentially damaged tissues or MSCs derived from bone marrow, the procurement of which requires a fairly morbid harvesting process with a low yield of stem cells. ASCs represent a healthy stem cell population that is easily procured by liposuction in large quantities. These cells can either be differentiated into SMCs in vitro or may be introduced directly into damaged tissues following in vitro expansion. Several studies have demonstrated that ASCs in an undifferentiated state that are injected into urological tissues remain viable and, over time, exhibit morphological and phenotypic evidence of SM incorporation and differentiation. In a pilot study, we isolated PLA cells from human adipose tissue by lipoaspiration, expanded them in culture, and labeled them with a fluorescent marker. These cells were then injected into the bladders or urethras of athymic or severe combined immunodeficiency (SCID) mice. Cells remained viable for up to 12 weeks following injection, and after 8 weeks demonstrated expression of early markers of SM differentiation (Jack et al., 2005). This study demonstrated the potential for ASCs to remain viable following delivery into urological tissues and to differentiate in vivo into SMCs, thus making them feasible, cost-effective, and a promising cell source for urinary tract reconstruction. These findings have recently been corroborated in another study where autologous ASCs obtained from the peri-gonadal fat pads of female Sprague–Dawley rats were injected into the proximal urethra and bladder following in vitro expansion and labeling with a fluorescent marker. In vivo differentiation into local recipient tissues was noted in 100% of the animals at the 4 and 6 week time points. They noted differentiation into smooth and striated muscle as well as vascular and adipose tissues, with differentiated SMCs expressing ASMA (Garcia et al., 2007). The capacity for ASCs to differentiate in vivo in the presence of other cells remains poorly understood. The mechanism by which this occurs is not currently known, with possibilities including cell plasticity and/or cell fusion of ASCs with surrounding cells. In a recent study attempting to delineate these mechanisms, differentiation of ASCs was induced by bladder SMCs both in vitro and in vivo. Chromosomal analysis was performed and revealed that cell fusion did not play a role in the induction of SM differentiation of ASCs. It is therefore felt that the in vivo SM differentiation of ASCs results from a combination of cell

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plasticity and influences from microenvironmental cues originating from surrounding cells and/or extracellular matrix interactions (Zhang et al., 2008). Regardless of the mechanism, ASCs may be differentiated in vitro into SMCs or may be introduced into the urinary tract in an undifferentiated state with subsequent induction of SM differentiation in vivo by host factors. Several studies incorporating ASCs into different facets of urinary tract reconstruction have been initiated and are currently in progress. Some of the preliminary results will be summarized in this chapter.

19.6.1 Bladder A number of injuries, both congenital and acquired, may affect the bladder resulting in the need for either augmentation cystoplasty or complete replacement of the bladder with or without urinary diversion. The mainstay of current bladder reconstruction in these cases generally involves the incorporation of intestinal segments into the urinary tract, which may be complicated by metabolic abnormalities, stone or mucus formation, as well as predisposition to urinary tract infections and even de novo malignancies (McDougal, 1992, Soergel et al., 2004). For this reason, there has been active investigation looking for alternative tissues and materials for bladder reconstruction (Bono and De Gresti, 1966, Cheng et al., 1994, Fujita, 1978, Gleeson and Griffith, 1992, Kelami et al., 1970, Kudish, 1957, Monsour et al., 1987, Neuhof, 1917, Probst et al., 1997, Rohrmann et al., 1996, Tsuji et al., 1961, Tsuji et al., 1967, Vaught et al., 1996). These numerous approaches have been fraught with limitations of structure, function, or biocompatibility. More recent studies incorporating cell-based therapies, however, have shown particular promise, avoiding the complications of fibrosis, scarring, and reduced reservoir volume that had been seen with the use of biodegradable scaffolds alone. Perhaps the most promising approach to date utilizes autologous urothelial and SM cells obtained from bladder biopsies that are seeded on to an allogenic acellular bladder matrix in the shape of a bladder. This construct was tested in a dog model where trigone-sparing cystectomy was performed; the bladder was either closed or replaced with the biodegradable scaffold that was seeded with autologous cells in one group and unseeded in another (Oberpenning et al., 1999, Yoo et al., 1998). Results from this study were promising, leading to further testing of this tissueengineered neobladder construct in a clinical trial involving seven patients with myelomeningocele and neurogenic bladders, in which scaffolds were seeded with autologous bladder epithelial and SM cells obtained from bladder biopsy. Results were promising with average follow-up of 5 years (Atala et al., 2006). The same group has recently published a follow-up study demonstrating minimal local and systemic side effects following the implantation of their cell-seeded neobladder in a canine model (Kwon et al., 2008).

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As promising as these advancements are, incorporation of SMCs derived from a healthy source may prove to have benefit over the use of SMCs obtained from a biopsy of the diseased organ. Although in earlier stages of research, ASCs have been used for the purpose of augmentation cystoplasty in a rat partial cystectomy model with great promise. ASCs from human lipoaspirate were differentiated in vitro into SMCs using SMIM. Nude rats then underwent partial cystectomy with either primary closure of the bladder or reconstruction with a three-dimensional, two-layered biodegradable bladder scaffold comprised of an electrospun sheet of poly(lactic-co-glycolic acid) (PLGA) under a porous PLGA sponge that was either seeded with the pre-differentiated SMCs or unseeded. Postoperative urodynamicsal evaluation with cystometry demonstrated maintenance of normal bladder capacity and compliance, and histological evaluation showed a normal three-dimensional tissue architecture and organization with persistence of an SM phenotype in those animals reconstructed with a cell-seeded neobladder. In contrast, those rats that underwent primary bladder closure or reconstruction with an unseeded scaffold demonstrated a time-dependent decrease in bladder capacity and compliance associated with poorly organized tissue regeneration (Jack et al., 2006). Although promising, further studies and clinical trials need to be undertaken in order to prove the efficacy, durability, lack of toxicity, and feasibility of using ASCs for human bladder reconstruction.

19.6.2 Urethra A number of urological disease processes – including inflammatory and traumatic strictures, congenital anomalies such as hypospadias, and malignancy – mandate the need for extensive urethral reconstruction. Although significant advances in management of these difficult problems have been made over the course of the past 50 years with the advent of microsurgical techniques, the ideal approach still eludes us, as evidenced by the multitude of potential reconstructive options. Over the past 20 years, tissue engineering strategies have been investigated as a means of urethral reconstruction. Numerous strategies have been employed using permanent and biodegradable, synthetic and biologically derived materials in various animal models (Chen et al., 1999, De Filippo et al., 2002, Italiano et al., 1998, Olsen et al., 1992, Sievert et al., 2000, Sievert et al., 2001). These experimental results in animal models have paved the road for several clinical studies that have shown promise for the use of an acellular collagen matrix for the purpose of urethral reconstruction using an onlay technique (El-Kassaby et al., 2003, El-Kassaby et al., 2008). A pre-clinical study evaluating the segmental replacement of urethra with a tubularized acellular collagen matrix seeded with autologous bladder epithelial and SM

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cells has also been reported in a canine model with good results in those animals receiving seeded grafts (Orabi, 2007). These data have laid the foundation for future directed studies evaluating alternative methods and strategies for segmental urethral reconstruction. Use of alternative cell sources for SM regeneration and/or alternative scaffold designs may provide improvements in the architecture and organization of regenerated SM, ultimately leading to better clinical outcomes for urethral reconstruction while eliminating the morbidity of and the need for tissue harvesting. Studies evaluating alternative scaffold designs and the use of ASCs for urethral reconstruction are currently underway in animal models.

19.6.3 Injectable therapies for treatment of stress urinary incontinence Although the pathophysiology of stress urinary incontinence (SUI) remains poorly understood, its etiology is likely multifactorial, resulting in part from loss of anatomic pelvic support and ischemic atrophy of pelvic floor musculature and in part from the ageing process, which induces loss of neuronal mass as well as atrophy of both the skeletal rhabdosphincter and the intrinsic smooth musculature of the urethra. The ultimate result is urethral dysfunction leading to loss of mucosal coaptation and failure to maintain resting urethral closure pressure. Following the general trend in surgery towards less invasive treatment modalities, minimally invasive treatments for SUI involving the injection of periurethral biologically derived or synthetic bulking agents to improve mucosal coaptation and increase urethral resistance have recently been described (Balmforth and Cardozo, 2003). These methods, though appealing owing to their minimally invasive nature, are fraught with limitations including lack of durability resulting in the need for repeat injections, as well as material antigenicity and migration (Echols et al., 2002, Kershen et al., 2002, Pannek et al., 2001). In addition, although these methods provide urethral bulking, they do not address the underlying pathophysiology of SUI. Tissue engineering strategies involving cell-based therapies have recently been introduced into the armamentarium of experimental treatments for SUI. These have included the periurethral injection of autologous chondrocytes, unprocessed fat, and skeletal muscle cells (Bent et al., 2001, Cannon et al., 2003, Chancellor et al., 2000, Furuta et al., 2007, Lee et al., 2003, Olson et al., 1998, Strasser et al., 2004, Trockman and Leach, 1995, Yokoyama et al., 2001). While the use of chondrocytes and fat was initially encouraging, long-term effects on continence parameters were much less promising. Recent technical advancements in muscle-derived stem cell selection and plating

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techniques have led to improvements in cell survival rates, propelling myoblast transfer therapy for treatment of SUI into the spotlight (Chancellor et al., 2000, Yokoyama et al., 2001). In an animal model of incontinence, investigators reported improved sphincter contractility and increased leak point pressures by urodynamic assessment following the periurethral injection of muscle-derived stem cells (Cannon et al., 2003, Chermansky et al., 2004, Lee et al., 2003). Furthermore, clinical trials have demonstrated the efficacy of transurethral ultrasound-guided injection of myoblasts into the rhabdosphincter with concomitant injection of fibroblasts into the urethral submucosa for treatment of SUI in humans (Strasser et al., 2004). These results have been validated in a prospective randomized trial comparing periurethral injection of myoblasts and fibroblasts with injection of collagen alone (Strasser et al., 2007). Multicenter trials incorporating larger numbers of patients with longer follow-up will need to be established in order to determine whether this treatment modality provides a durable response that may be standardized between different centers. While these studies demonstrate exciting and promising results and show the potential for autologous cell transplantation in the treatment of SUI, they primarily target the regeneration of skeletal muscle within the rhabdosphincter with fibroblast proliferation accounting for submucosal bulking. This strategy does not lead to regeneration of the urethral SM layer, which accounts for approximately 50% of the intrinsic continence mechanism, and the loss of which contributes significantly to intrinsic sphincter deficiency (Petros and Ulmsten, 1998). In this regard, additional strategies for the regeneration of the intrinsic urethral SM layer may further target the underlying pathophysiology of SUI and provide additional benefit in its treatment. We have demonstrated that ASCs injected periurethrally exhibit in vivo differentiation into SM and therefore show promise as a potential cell source for the regeneration of SM for the purpose of SUI treatment (Jack et al., 2005). In this study, cell dispersion and migration were noted 4 weeks following injection. In order to address this limitation, additional tissue engineering strategies involving the co-injection of a biodegradable scaffold as a temporary support structure for cell attachment, proliferation, and SM regeneration were conceived. In one study, periurethral co-injection of human SM cells differentiated from ASCs (SM-ASCs) with PLGA microspheres was compared with injection of saline alone, SM-ASCs alone, human fibroblasts alone, fibroblasts co-injected with PLGA microspheres, and PLGA microspheres alone following transabdominal urethrolysis in immunocompromised nude rats (Zeng, 2007). Animals underwent in vivo urodynamic assessment of continence parameters as well as isometric studies and histological analysis of harvested urethral tissues following killing at 2, 6, or 12 weeks post-injection. A durable and reproducible decrease in abdominal leak point pressure (ALPP) and retrograde urethral

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perfusion pressure (RUPP) was noted in those animals that underwent transabdominal urethrolysis with either no injection or with injection of saline alone. In animals injected with cells alone, progressive cell migration was seen over time. The co-injection of PLGA microspheres led to an immediate bulking effect with resulting improvements in continence parameters and facilitated the long-term localization of cells to the periurethral area. Long-term improvements in ALPP and RUPP were only seen in the group of animals co-injected with SM-ASCs and PLGA microspheres. In addition, isometric urethral contraction in response to pharmacological agents was only improved in those animals that were injected with SM-ASCs. Histologically, injected SMCs maintained both early and late markers of SM differentiation at the 12 week timepoint. This study validates the potential for tissue engineering strategies incorporating ASCs for the treatment of SUI.

19.6.4 Erectile dysfunction and male infertility The mainstay of treatment for erectile dysfunction has classically involved the use of phosphodiesterase-5 inhibitors, intracavernous or intraurethral treatments with erectogenic substances, vacuum devices, or surgical placement of penile prostheses. In cases of infertility caused by primary testicular failure, management has been supportive with the use of testosterone replacement therapy with minimal options available for inducing spermatogenesis. Recent advancements in stem cell research have raised the possibility of restoring erectile function through restoration and/or regeneration of corpora cavernosal SM and cavernous nerves. In addition, gonadal function may potentially be restored through regeneration of Leydig, Sertoli, and germ cells. For the treatment of erectile dysfunction, the use of MSCs derived from bone marrow transduced with endothelial nitric oxide synthase (eNOS) has been shown to improve the erectile function of aged rats and even reverse age-associated erectile dysfunction (Bivalacqua et al., 2007, Deng et al., 2003). Additionally, the use of embryonic stem cells transfected with brainderived neurotrophic factor (BDNF) has been shown to restore erectile function in a cavernous nerve injury model of neurogenic erectile dysfunction (Bochinski et al., 2004). As an adjunct, biodegradable scaffolds incorporating neurotrophic factors have been employed to facilitate nerve regeneration following injury to the cavernous nerves (May et al., 2005). Cell-based therapies targeting corpora cavernosal SM and endothelial regeneration using autologous SMCs or bone marrow-derived MSCs in animal models have also been promising (Kwon et al., 2002, Song et al., 2007). Realizing the vast potential for the use of ASCs in treatment of erectile dysfunction, Tom Lue and associates at the University of California

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San Francisco have reported on the successful use of ASCs for the treatment of neurogenic erectile dysfunction in rats (Lin et al., 2008). With regards to male infertility, there have been several landmark studies since 2003, when it was reported that embryonic stem cells could form functional male germ cells in vitro (Toyooka et al., 2003). In 2006, the first evidence was presented that germ cells derived from embryonic stem cells could form functional male gametes with the capability of generating offspring mice following transplantation (Nayernia et al., 2006b). The same group later presented evidence that functional germ cells could be differentiated in vitro from murine as well as human bone marrow-derived MSCs (Drusenheimer et al., 2007, Nayernia et al., 2006a). In a mouse infertility model, MSCs transplanted into the testes seem to differentiate into germ cells as well as Sertoli and Leydig cells (Lue et al., 2007). This same group is currently investigating whether ASCs may differentiate in a similar manner, providing a novel treatment approach for hypogonadism as well as male infertility.

19.7

Future trends

Adipose-derived stem cells represent an abundant, easily procured source of healthy, multipotent adult MSCs that show great promise and utility for the purpose of urinary tract reconstruction. Whether these cells represent a circulating pool of progenitor cells originating within the bone marrow stroma or a distinct population of cells originating within adipose tissue remains debatable, as evidenced by the ongoing controversy surrounding the appropriate nomenclature for these cells. Although their molecular characterization has lagged behind that of bone marrow derived MSCs, as their utility in the regeneration of multiple mesenchymally derived tissues is realized, their specific characterization will become paramount in providing standardization of isolation and differentiation processes between different laboratories. This realization will hopefully spur further research into ASC molecular characterization, thereby also providing greater insight into the origin of these cells. In addition, the ability of certain stem cell populations to migrate and home in on sights of tissue injury is currently being widely investigated by numerous groups. Although this topic is beyond the scope of this chapter, the process seems to be coordinated by numerous cellular interactions with soluble homing factors and other molecular mechanisms that remain to be elucidated. Research into these mechanisms as well as into various methods for the tracking of these cells will not only be critical in our understanding of stem cell biology, but may also provide insight into new and innovative approaches for the delivery of stem cell therapies to tissues of interest both within the urinary tract and without.

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Numerous advances have been made in the regeneration and restoration of urinary tract tissues using tissue engineering strategies. Several studies have been pivotal in defining the potential for ASCs to differentiate into SMCs and vascular and neurological structures, both in vitro and in vivo, thereby opening the door for the use of ASCs in cell-based therapies targeting the urinary tract. Numerous factors including extracellular matrix interactions and growth factor/cytokine effects are known to influence the differentiation capacity of ASCs. Studies to further elucidate the optimal differentiation conditions for ASCs should contribute to the wider acceptance of this cell source for tissue engineering applications. In addition, the process by which ASCs can be induced to differentiate into different mesenchymal tissues in the presence of already differentiated tissues remains to be clarified. A better understanding of this process may facilitate improved differentiation of ASCs in an in vitro or in vivo setting, providing further optimization of tissue engineering strategies. The advantages of ASCs as a source for cell-based therapies have recently been realized. Investigation of their use is currently under way for the treatment of erectile dysfunction, male infertility, SUI, and for bladder and urethral reconstruction. Tissue engineering strategies for the purpose of bladder and urethral reconstruction as well as for treatment of SUI have resulted in human clinical trials that have demonstrated exciting results with great promise. With respect to bladder and urethral reconstruction, these approaches have the theoretical limitation of utilizing autologous cells derived from diseased tissues or organs. With respect to injectable SUI treatments, current strategies target skeletal muscle regeneration of the urethral rhabdosphincter, which may not fully address the pathophysiology of SUI, namely the atrophy of urethral smooth musculature that plays a significant role in intrinsic sphincter deficiency. In conjunction with advances in scaffold design, improvements in delivery of growth factors that facilitate SM differentiation, incorporation of techniques to apply mechanical strain in order to improve SM differentiation, and the incorporation of stem cellderived neurological and vascular structures, the use of ASC-based cellular therapy in combination with current strategies will hopefully provide added advantages that will make the use of ASCs standard for urinary tract reconstruction in the future.

19.8

References

atala, a., bauer, s. b., soker, s., yoo, j. j. and retik, a. b. (2006) Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet, 367, 1241–6. balmforth, j. and cardozo, l. d. (2003) Trends toward less invasive treatment of female stress urinary incontinence. Urology, 62, 52–60.

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tsuji, i., ishida, h. and fujieda, j. (1961) Experimental cystoplasty using preserved bladder graft. J Urol, 85, 42–4. tsuji, i., shiraishi, y., kassai, t., kunishima, k., orikasa, s. and abe, n. (1967) Further experimental investigations on bladder reconstruction without using the intestine. J Urol, 97, 1021–8. urbich, c. and dimmeler, s. (2004) Endothelial progenitor cells functional characterization. Trends Cardiovasc Med, 14, 318–22. van der loop, f. t., schaart, g., timmer, e. d., ramaekers, f. c. and van eys, g. j. (1996) Smoothelin, a novel cytoskeletal protein specific for smooth muscle cells. J Cell Biol, 134, 401–11. van harmelen, v., rohrig, k. and hauner, h. (2004) Comparison of proliferation and differentiation capacity of human adipocyte precursor cells from the omental and subcutaneous adipose tissue depot of obese subjects. Metabolism, 53, 632–7. vaught, j. d., kropp, b. p., sawyer, b. d., rippy, m. k., badylak, s. f., shannon, h. e. and thor, k. b. (1996) Detrusor regeneration in the rat using porcine small intestinal submucosal grafts: functional innervation and receptor expression. J Urol, 155, 374–8. wagner, w., wein, f., seckinger, a., frankhauser, m., wirkner, u., krause, u., blake, j., schwager, c., eckstein, v., ansorge, w. and ho, a. d. (2005) Comparative characteristics of mesenchymal stem cells from human bone marrow, adipose tissue, and umbilical cord blood. Exp Hematol, 33, 1402–16. williams, b. (1998) Mechanical influences on vascular smooth muscle cell function. J Hypertens, 16, 1921–9. williams, s. k., wang, t. f., castrillo, r. and jarrell, b. e. (1994) Liposuction-derived human fat used for vascular graft sodding contains endothelial cells and not mesothelial cells as the major cell type. J Vasc Surg, 19, 916–23. winter, a., breit, s., parsch, d., benz, k., steck, e., hauner, h., weber, r. m., ewerbeck, v. and richter, w. (2003) Cartilage-like gene expression in differentiated human stem cell spheroids: a comparison of bone marrow-derived and adipose tissuederived stromal cells. Arthritis Rheum, 48, 418–29. wolf, n. s., penn, p. e., rao, d. and mckee, m. d. (2003) Intraclonal plasticity for bone, smooth muscle, and adipocyte lineages in bone marrow stroma fibroblastoid cells. Exp Cell Res, 290, 346–57. wright, j. t. and hausman, g. j. (1990a) Adipose tissue development in the fetal pig examined using monoclonal antibodies. J Anim Sci, 68, 1170–5. wright, j. t. and hausman, g. j. (1990b) Monoclonal antibodies against cell surface antigens expressed during porcine adipocyte differentiation. Int J Obes, 14, 395–409. yang, y., relan, n. k., przywara, d. a. and schuger, l. (1999) Embryonic mesenchymal cells share the potential for smooth muscle differentiation: myogenesis is controlled by the cell’s shape. Development, 126, 3027–33. yokoyama, t., yoshimura, n., dhir, r., qu, z., fraser, m. o., kumon, h., de groat, w. c., huard, j. and chancellor, m. b. (2001) Persistence and survival of autologous muscle derived cells versus bovine collagen as potential treatment of stress urinary incontinence. J Urol, 165, 271–6. yoo, j. j., meng, j., oberpenning, f. and atala, a. (1998) Bladder augmentation using allogenic bladder submucosa seeded with cells. Urology, 51, 221–5. yoshimura, k., shigeura, t., matsumoto, d., sato, t., takaki, y., aiba-kojima, e., sato, k., inoue, k., nagase, t., koshima, i. and gonda, k. (2006) Characterization of

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freshly isolated and cultured cells derived from the fatty and fluid portions of liposuction aspirates. J Cell Physiol, 208, 64–76. young, c., jarrell, b. e., hoying, j. b. and williams, s. k. (1992) A porcine model for adipose tissue-derived endothelial cell transplantation. Cell Transplant, 1, 293–8. zeng, x., jack, g. s., zhang, r., leung, j., wu, b. and rodriguez, l. v. (2007) Fibroblasts and adipose-derived stem cells in the treatment of stress urinary incontinence: comparison of functional outcomes: Presented at the American Urological Association Annual Meeting 2007; abstract. zhang, r., jack, g. s., rao, p. n. and rodriguez, l. v. (2008) Cell fusion-independent smooth muscle differentiation of human adipose derived cells induced by smooth muscle environment: Presented at the American Urological Association Annual Meeting 2008; UCLA Medical Center Los Angeles, CA; abstract. zuk, p. a., zhu, m., ashjian, p., de ugarte, d. a., huang, j. i., mizuno, h., alfonso, z. c., fraser, j. k., benhaim, p. and hedrick, m. h. (2002) Human adipose tissue is a source of multipotent stem cells. Mol Biol Cell, 13, 4279–95. zuk, p. a., zhu, m., mizuno, h., huang, j., futrell, j. w., katz, a. j., benhaim, p., lorenz, h. p. and hedrick, m. h. (2001) Multilineage cells from human adipose tissue: implications for cell-based therapies. Tissue Eng, 7, 211–28. zvaifler, n. j., marinova-mutafchieva, l., adams, g., edwards, c. j., moss, j., burger, j. a. and maini, r. n. (2000) Mesenchymal precursor cells in the blood of normal individuals. Arthritis Res, 2, 477–88.

20 Regenerative medicine of the urinary sphincter via an endoscopic approach M. C. S M A L D O N E, University of Pittsburgh School of Medicine, USA; and M. B. C H A N C E L L O R, William Beaumont Hospital, USA

Abstract: In anatomical and functional studies of the human and animal urethra, the middle urethral-contained rhabdosphincter is critical for maintaining continence. Transplanted stem cells may have the ability to undergo self-renewal and multipotent differentiation, leading to sphincter regeneration. In addition, such cells may release, or be engineered to release, neurotrophins with subsequent paracrine recruitment of endogenous host cells to promote concomitantly a regenerative response of nerve-integrated muscle. Cell-based therapies are most often associated with the use of autologous multipotent stem cells, such as the bone marrow stromal cells. However, harvesting bone marrow stromal stem cells is difficult and painful, and may yield low numbers of stem cells upon processing. In contrast, alternative autologous adult stem cells such as muscle-derived stem cells and adipose-derived stem cells can be easily obtained in large quantities and with minimal discomfort. This chapter aims to discuss the following: the neurophysiology of stress urinary incontinence (highlighting the importance of the middle urethra); current injectable cell sources for endoscopic treatment; and the potential of muscle-derived stem cells for the delivery of neurotrophic factors. Key words: stem cells, muscle, urinary incontinence, urethral sphincter.

20.1

Introduction

There are over 200 million people worldwide with incontinence, a condition that is associated with an adverse social impact and a reduced quality of life (Corcos et al., 2002, Norton and Brubaker, 2006). Stress urinary incontinence (SUI) has been reported as the most common type of urinary incontinence (Hampel et al., 1997). Risk factors for developing SUI include increasing parity, age, and obesity (Bump and Norton, 1998). Injury during childbirth to pelvic floor musculature, connective tissue, and nervous structures appears to be the most important risk factor for urinary incontinence in later life (Peschers et al., 1996, Thom et al., 1997, Meyer et al., 1998, Sampselle et al., 1998). 422

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SUI can be grouped into two major categories: urethral hypermobility and intrinsic sphincter deficiency (ISD) (McGuire et al., 1976). Urethral hypermobility, or loss of bladder neck support, results in a lack of intraabdominal pressure transmission to the proximal urethra. In contrast, ISD is characterized by malfunction of the urethral closure mechanism. However, the grouping of patients into dichotomous categories has not translated into diagnostic and therapeutic improvements (Blaivas, 1987). Most researchers now believe that SUI varies between the extremes of urethral hypermobility and ISD, with most patients having elements of both disorders (Kayigil et al., 1999). The role of pharmacotherapy for SUI, including α-adrenoceptor agonists, has been disappointing (Radley et al., 2001). Duloxetine, a selective serotonin and norepinephrine reuptake inhibitor, is available in Europe but it was not approved by the Food and Drug Administration (FDA) in the United States for the indication of SUI (Chancellor, 2004). The use of injectable bulking agents including polytetrafluoroethylene (Herschorn and Glazer, 2000), bovine collagen (Sakamoto et al., 2007), silicone particles (Maher et al., 2005), carbon beads (Chrouser et al., 2004), and autologous ear chondrocytes (Bent et al., 2001) has yielded short-term success in the treatment of SUI. However, use of bulking agents has resulted in chronic inflammatory reactions, foreign body giant cell responses, periurethral abscess, particle migration, erosion of the urinary bladder or the urethra, obstruction of the lower urinary tract with urinary retention, severe voiding dysfunction, and pulmonary embolism (Kiilholma et al., 1993, Papa Petros, 1994, McKinney et al., 1995, Sweat and Lightner, 1999). The potential of stem cell therapy for the regenerative repair of the deficient rhabdosphincter is currently at the forefront of incontinence research (Lee et al., 2003). Overall, the aim of stem cell therapy is to replace, repair, or enhance the biological function of damaged tissue or organs. There are two general types of stem cells potentially useful for therapeutic treatment: embryonic stem cells (ESCs) and adult stem cells. Although theoretically appealing, the practical use of ESCs is limited due to problems of cell regulation and ethical considerations (Edwards, 2007). In contrast, adult stem cells have no significant ethical issues related to their use. We envision that in the near future treatment of SUI will involve a routine visit to a urologist or urogynecologist’s office where a muscle biopsy using a small-caliber needle biopsy device will be performed. The biopsy is preserved and shipped to a central approved stem cell facility for processing, where a number of muscle-derived stem cells (MDSCs) are isolated, prepared, and stored. Within a period of weeks, the MDSCs are shipped back to the doctor’s office and, using a cystoscopy or spinal needle, the cells are injected into the patient’s urethral sphincter under local anesthesia (Fig. 20.1).

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20.1 Diagram showing autologous stem cell injection therapy for stress urinary incontinence. Autologous stem cells are obtained with a biopsy of tissue, the cells are dissociated and expanded in culture, and the expanded cells are implanted into the same host. From Furuta et al. (2007a), with permission from John Wiley & Sons, Inc.

20.2

Neurophysiology of stress urinary incontinence

All of the urethral muscles, pelvic floor muscles, and surrounding connective tissues contribute to urethral resting tone. When intra-abdominal pressures are increased by coughing, sneezing, or exercise, the urethra is passively or actively closed to prevent urinary leakage. This passive urethral closure mechanism is well understood. Contraction of the pelvic floor muscles (levator ani muscles) pulls the vagina forward toward the pubic symphysis, creating a backstop for the urinary tract. This stable backstop compresses the two walls of the urethra, preventing urinary leakage during elevation of intra-abdominal pressures (DeLancey, 1994). In addition, the position of the bladder neck is important for pressure transmission to the properly positioned structures during increased intra-abdominal pressure to remain equal. Descent of the structures causes a pressure gradient that can result in urinary leakage. The active urethral closure mechanism that maintains urinary continence during elevation of intra-abdominal pressures is currently still under inves-

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tigation. The external urethral sphincter (EUS), commonly referred to as the rhabdosphincter, is composed of both type I and type II striated muscle fibers located in the middle urethra (Schroder and Reske-Nielsen, 1983). In addition, urethral smooth muscles are deposited in longitudinal and circular layers. The urethral muscles are controlled by three sets of peripheral nerves: sacral parasympathetic nerves (pelvic nerves), thoracolumbar sympathetic nerves (hypogastric nerves), and sacral somatic nerves (pudendal nerves) (Fig. 20.2). Sympathetic preganglionic pathways emerge from the thoracolumbar spinal cord, pass through the sympathetic chain ganglia (SCG) and the inferior splanchnic nerves (ISN), and on to the inferior mesenteric ganglia (IMG). Preganglionic and postganglionic sympathetic axons then travel through the hypogastric nerves to the pelvic plexus and the urogenital

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20.2 Diagram showing the sympathetic, parasympathetic, and somatic innervation of the urogenital tract. Abbreviations: IMG, inferior mesenteric ganglia; SCG, sympathetic chain ganglia; DRG, dorsal root ganglia; ISN, inferior splanchnic nerves; EUS, external urethral sphincter. From Furuta et al. (2007a), with permission from John Wiley & Sons, Inc.

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organs. Parasympathetic preganglionic axons that originate in the sacral spinal cord pass through the pelvic nerves to ganglion cells located in the pelvic plexus and then to distal ganglia in the target urogenital organs. Sacral somatic pathways are contained in the pudendal nerves, which provide innervation to the EUS. The pudendal and pelvic nerves also receive postganglionic axons from the caudal SCG. These three sets of nerves contain afferent axons from the lumbosacral dorsal root ganglia (DRG) (de Groat, 2006). Recent studies have reported that the EUS can be activated voluntarily or by a reflex mechanism elicited by bladder distension (Chancellor et al., 2005). Nerve tracts from the central nervous system terminate at Onuf’s nucleus in the sacral spinal cord and synapse with the pudendal nerves. Serotonin and norepinephrine are two key neurotransmitters that stimulate the proximal end of the pudendal nerves to control EUS contraction (Thor and Katofiasc, 1995, Chancellor et al., 2005, Kaiho et al., 2007). The nerve-mediated active urethral closure mechanism that maintains urinary continence during elevation of abdominal pressures may be divided into two areas of control: central nervous control passing through Onuf’s nucleus during sneezing or coughing (Kamo et al., 2003, Kamo et al., 2006), and the bladder-to-urethral spinal reflex during laughing, exercise, or lifting heavy objects (de Groat, 1998, Kamo et al., 2004). In an elegant series of experiments, Kamo et al. (de Groat, 1998, Kamo et al., 2003, Kamo et al., 2004, Kamo et al., 2006) demonstrated that an increase in middle urethral pressure during sneezing was caused not only by passive transmission of increased abdominal pressure but also by active reflex contractions of the EUS and pelvic floor muscles. In addition, they were able to show that this sneeze-induced continence reflex in the middle urethra is impaired in a rat SUI model induced by vaginal distension. In contrast, the increase in proximal and distal urethral pressure during sneezing was dependent upon increases in intravesicular and/or intra-abdominal pressure (Fig. 20.3). Another key observation was that passive elevation of intravesicular pressure in spinal cord-transected rats elicited both pelvic afferent nerve-mediated contractile reflexes and bladder-to-urethral spinal reflexes in the middle-to-proximal urethra mediated by the activation of pelvic nerves (Fig. 20.4). In anatomical reports, Strasser et al. (2000a) showed that hypogastric and pelvic nerves predominantly innervated and regulated the proximal urethra, whereas stimulation of pudendal nerves led to contraction of the middleto-distal urethra. In addition, a decrease in the number of striated muscle cells was reported in conjunction with an age-dependent increase of apoptosis of the striated muscle fibers of the EUS (Strasser et al., 2000b). These results suggest that the middle urethra and EUS are critical for maintaining continence and represent a primary focus in the management of SUI.

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20.3

Stem cell source for the injection therapy of stress urinary incontinence

Cell-based therapies and tissue engineering are most often associated with the use of autologous multipotent stem cells. One commonly described source of such cells is the bone marrow stroma. The bone marrow compartment contains several cell populations including mesenchymal stem cells (MSCs) that are capable of differentiating into adipogenic (Pittenger et al., 1999), osteogenic (Pittenger et al., 1999), chondrogenic (Ferrari et al., 1998), and myogenic cells (Prockop, 1997, Dezawa et al., 2005). However, autologous bone marrow procurement has significant inherent limitations including painful procurement procedures that frequently require general or spinal

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20.4 Changes in urethral pressure responses at different sites in the urethra (2.5 mm steps from the urethral orifice) measured by a microtip transducer catheter during increments in intravesicular pressure to 20, 40, or 60 cmH2O. Note that intravesicular pressure elevation-induced urethral response was the biggest in the urethra at 12.5–15 mm, whereas the baseline was highest in the urethra at 10–12.5 mm from the orifice. From Kamo et al. (2004), with permission from the American Physiological Society.

anesthesia, and often yield a low number of MSCs upon processing (Pittenger et al., 1999). As an alternative source of autologous adult stem cells, MDSCs and adipose-derived stem cells (ADSCs) are advantageous because they can be easily obtained in large quantities under local anesthesia. Muscle-derived cell therapy, often referred to as myoblast transfer therapy, has in the past been hindered by numerous limitations including poor survival of the injected cells. Selection of specific stem cells from the pool of remaining muscle-derived cell populations, through the use of techniques such as pre-plating, has led to improved cell survival rates following transplantation (Qu et al., 1998). Such observations have since led to extensive investigation into the developmental origins of skeletal muscle progenitor cells and the functional heterogeneity displayed by various skeletal musclederived cell populations (Seale et al., 2001, Jankowski et al., 2002). While difficult to identify in vitro through expression of specific ‘marker’ proteins,

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which are often in flux or may be down-regulated quickly following placement in culture, MDSCs display a remarkable regenerative capacity when compared with the commonly recognized and abundant striated muscle precursor ‘myoblast’ cells. The total regenerative response elicited through the transplantation of such cells – including their survival, engraftment, induction of host tissue repair, and ability to restore functionality – is elegantly displayed in a recent report comparing the fate of MDSCs and myoblasts in a myocardial infarct model (Oshima et al., 2005). Therefore, not all musclederived cells are created equal and care should be taken when considering and comparing reports originating from different laboratories. In our ongoing research, MDSCs display an improved transplantation capacity with the ability to undergo long-term proliferation, self-renewal, and multipotent differentiation, including differentiation toward endothelial and neuronal lineages (Lee et al., 2000, Qu-Petersen et al., 2002). MDSC injection therapy offers several advantages over conventional treatments for SUI. The use of cells that are derived from the incontinent patient (autologous cell transplantation) will not cause an immunogenic or allergic reaction and therefore may persist longer than injected foreign substances such as collagen (Yokoyama et al., 2001b, Lee et al., 2004). MDSCs are uniquely different to fibroblasts and smooth muscle cells since MDSCs will fuse to form post-mitotic multinucleated myotubes (Fig. 20.5). This limits the persistent expansion and risk of obstruction that may occur with other cell sources such as fibroblasts (Kwon et al., 2006). Finally, MDSCs form myotubes and myofibers that become innervated into the host muscle (Fig. 20.6). Therefore, not only can they serve as a bulking agent, but also they

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20.5 Human muscle-derived stem cells proliferating in growth culture medium. (a) Multinucleated myotube structures formed following placement in differentiation medium. (b) Immunofluorescence of myogenic-specific proteins desmin (gray) and myosin heavy chain (white). From Furuta et al. (2007b), with permission from Current Medicine Group, LLC.

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20.6 Hematoxylin–eosin staining revealed (a) atrophic proximal urethral sphincter in saline group at 4 weeks compared with (b) normal, uninjured urethra. (c) and (d) Muscle-derived cells (MDCs) (1 × 106) injected into denervated sphincter led to increased dorsolateral striated muscle masses with variable fiber orientation at injection sites. (e) and (f) Fibroblasts (1 × 106) injected into denervated sphincter led to increased dorsolateral connective tissue masses at injection sites. Images taken with ×10 ((a), (b), (c), and (e)) and ×20 ((d) and (f)) objectives. From Kwon et al. (2006), with permission from Elsevier.

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are physiologically capable of improving urethral sphincter function (Chancellor et al., 2000, Yokoyama et al., 2001a, Huard et al., 2002). The feasibility of this concept was first demonstrated in rat models of SUI (Cannon et al., 2003, Lee et al., 2003). Chermansky et al. (2004) showed that MDSCs had integrated within the striated muscle layer of the cauterized middle urethra 4 weeks after injection (Fig. 20.7). In addition, the striated muscle layer of the MDSC-injected urethra was contiguous with an increase in nervous tissue when compared with those of the cauterized urethra injected with only saline solution (Fig. 20.8). These results suggest that MDSCs may have the capacity for multipotent differentiation in the host urethral tissue or have the capacity to elicit a paracrine effect resulting in a more complete regenerative muscle–nerve healing response. In

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20.7 Histological findings of cauterized mid-urethra 4 weeks after Hanks’ balanced salt solution (HBSS) or MDC injection. (a) Hematoxylin–eosin stain of cauterized mid-urethra injected with HBSS (×400). Arrow points to disrupted striated muscle layer. (b) Hematoxylin-eosin stain of cauterized mid-urethra injected with MDCs (×400). Arrow points to intact striated muscle layer. (c) LacZ stain of cauterized mid-urethra injected with MDCs (×400). Arrows point to β-galactosidase-expressing MDCs, situated within striated muscle layer of mid-urethra. From Chermansky et al. (2004), with permission from Elsevier.

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20.8 Differences in striated muscle layer and innervation of cauterized mid-urethra. (a) Fast myosin heavy chain stain of cauterized midurethra 4 weeks after HBSS injection (×400). Arrows point to disrupted striated muscle layer. (b) Fast myosin heavy chain stain of cauterized mid-urethra 4 weeks after MDC injection (×400). Arrow points to intact striated muscle layer. (c) Protein gene product (PGP) 9.5 stain of cauterized mid-urethra 4 weeks after HBSS injection (×400). Arrows point to few nerve fibers present. (d) PGP 9.5 stain of cauterized midurethra 4 weeks after MDC injection (×400). Arrows point to many nerve fibers present. From Chermansky et al. (2004), with permission from Elsevier.

addition, the increase in leak point pressure (LPP) seen in the groups injected with MDSCs was significant when compared with the salineinjected cauterized rats. Importantly, the difference in LPP both 4 and 6 weeks after MDSC injection was not significant when compared with the uncauterized control rats (Fig. 20.9). In a recent report, Kwon et al. (2006) compared MDSCs and fibroblasts with regard to their potential for restoration of urethral function following injection. Using LPP for functional comparison, the short-term experiment revealed no significant difference between MDSCs and fibroblasts, or a combination of both, when the cell dosage was equal across the groups. However, when the dosage was varied by two 10-fold increases, only a high dose of fibroblast injection led to urinary retention. Importantly, even high

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doses of MDSCs did not result in such adverse events. These results suggest that fibroblasts may be producing a bulking effect, as evidenced by the increase in LPP, but also may make the tissue less compliant. It has been well documented that the volume and location of synthetic substance injections are extremely critical in achieving a low incidence of adverse events such as retention (Beckingham et al., 1992). Studies supporting the potential of ADSCs are also emerging. Zuk et al. (2001) described the differentiation of ADSCs in vitro into adipogenic, myogenic, and osteogenic cells in the presence of lineage-specific induction factors. In addition, ADSCs exhibited the functional ability to contract and relax in direct response to pharmacological agents (Rodriguez et al., 2006). ADSCs may also represent an alternative stem cell source for the treatment of SUI (Jack et al., 2005). The feasibility of ADSC use has been suggested through reports of improvements in LPP and urethral function in a rat model of SUI when animals were injected with ADSCs in conjunction with biodegradable microbeads as a carrier (Zeng et al., 2006). By providing a potential cost-effective source for genitourinary reconstruction, cell therapies using MDSCs and ADSCs are emerging as a promising technology for the treatment of SUI.

20.4

Role of muscle-derived stem cells in the delivery of neurotrophic factors

The use of growth factor proteins to promote healing is severely hindered by the difficulty of ensuring their delivery to the injured site, their short biological half-lives, and the rapid clearance of these molecules from the

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bloodstream (Robbins et al., 1998). Various growth factors appear to regulate skeletal myoblast proliferation and differentiation, play a role in different stages of muscle regeneration, and enhance the healing process (Florini et al., 1991). It is intriguing that adult knockout mice expressing a neutralizing antibody against nerve growth factor (NGF) display a severely reduced muscle mass (Ruberti et al., 2000). In addition to acting as a targetderived factor for developing neurons, NGF has an autocrine effect on myoblast proliferation and fusion (Rende et al., 2000). In fact, myoblasts express low-affinity p75 NGF receptors (Reddypalli et al., 2005) and highaffinity tyrosine kinase A (trkA) NGF receptors (Wheeler et al., 1998). The presence of NGF receptors and exposure of myoblasts to NGF resulted in upregulation of anti-apoptosis/pro-survival proteins. This suggests that NGF mediates survival of myoblasts prior to differentiation and is important for muscle fiber development. Since Schwann cells can survive inside the intramuscular nerve trunks of denervated skeletal muscles for a 25 month period without axonal contact in rats (Dedkov et al., 2002), NGF release by transplanted MDSCs may also promote axonal regeneration and functional recovery after nerve injury. It stands to reason that transplanted MDSCs that have the capacity to undergo self-renewal and multipotent differentiation, as well as release growth factors such as NGF, may promote a more complete response due to both autocrine and paracrine effects, leading to both a muscle and integrated nerve regenerative response of donor transplanted and host cells. In a number of tissues, the development and survival of sympathetic neurons are dependent upon the presence of target-derived neurotrophins, of which the best characterized is NGF (Shooter, 2001). Sympathetic activation of adipocytic β-adrenoceptors induces lypolysis and a decrease in the number of adipocytes. This process is responsible for a loss of body weight during hibernation (Shi et al., 2005). A recent study has demonstrated that NGF is synthesized and released by white adipose tissue with the expression of p75 and trkA NGF receptors in adipocytes (Peeraully et al., 2004). This suggests that ADSCs may also have the potential of releasing NGF and regenerating urethral malfunction in patients with SUI.

20.5

Injection technique

20.5.1 Percutaneous needle muscle biopsy (Lacomis, 2000) After informed consent has been obtained, the patient should be positioned in the supine position on a padded examination table. In order to aid in identifying the lateral muscle of the thigh (vastus lateralis muscle), the patient is instructed to contract their thigh. The proper biopsy site is located

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midway between the patella and the greater trochanter, at the anterior border of the iliotibial band. After the proper site is identified, the area should be cleaned with an antiseptic soap followed by an alcohol swab. Antibiotic prophylaxis for this procedure is rarely indicated. After cleaning, the skin and fascia overlying the muscle are anesthetized with approximately 1.5 cc of a local anesthetic. Following local anesthesia administration, the biopsy site and surrounding area is prepped with betadine solution and a sterile drape is applied. When the injected area over the biopsy site is numb, a half-centimeter incision is made with a ⱅ11 scalpel to the level of the fascia. At this point the patient is instructed to straighten and completely relax their thigh, and the muscle biopsy (60–100 milligrams) is performed with a 10 gauge QuickCore® disposable biopsy needle (Cook, Bloomington, IN, USA). After the biopsy needle is withdrawn, pressure is applied to the site to prevent bleeding, and the incision is closed with a steri-strip bandage followed by a pressure dressing. The subject is advised against all vigorous activity during the first 48 hours post-biopsy and is also advised not to shower or get the incision wet during this time period. Although the patient may feel moderate pressure during the biopsy, post-procedure discomfort is usually related to localized aching that rarely requires narcotic analgesia.

20.5.2 Endoscopic injection technique (Appell and Winters, 2007) After the patient has been counseled regarding the purpose of the procedure, it is important to ensure that the patient does not have an active urinary tract infection before cystourethroscopy, because of the possibility of exacerbating the infection by instrumentation of the urinary tract. In the female patient, 5–10 mL of lubricant–anesthetic jelly should be instilled into the urethra before the procedure. It is important to ensure sterility and obtain adequate urethral anesthesia for diagnostic cystourethroscopy. A perioperative dose of a prophylactic antibiotic should be administered, and except under extraneous circumstances routine post-operative antibiotics are unnecessary. Inspection of the female urethra is easily performed by inserting the endoscope under direct vision into the urethral meatus and by directing the instrument cephalad toward the umbilicus. The injections can be performed transurethrally through a 10 mm needle placed directly through a cystourethroscope or periurethrally with a spinal needle or specialty injector injected percutaneously and positioned in the urethral tissues while observing the manipulation endoscopically per urethra. The cause of the incontinence, the tissue at the injection site, and the plane of delivery of the injectable substance will affect the treatment results. Regardless of technique used, nearly

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every patient can be injected with the use of local anesthesia. In our clinical experience, use of a 10 mm injection needle facilitates deeper delivery into the external sphincter. Precise localization of the site of deposition is essential to ensure an optimal response. The periurethral approach decreases bleeding complications, which can hamper visualization and extrusion of the injected material. In a prospective randomized comparison of the transurethral and periurethral approaches with conventional bulking agent injection in women no differences in efficacy were seen, but urinary retention rates and the volume of material injected was higher in the periurethral group (Schulz et al., 2004). In the transurethral endoscopic approach, a 0°, 12°, or 30° lens can be used. The endoscope is placed at the mid-urethra, and the needle is advanced to near the mid-urethra and injected into the deep muscle layer at three or more spots. This differs from the submucosal collagen injection technique. The transurethral injection is guided by location and approximately 4 mL of material injected (containing 2–50 × 106 MDSCs). The volume and number of cells injected is currently being studied, and procedure success should not be determined by the amount of urethral coaptation. In our experience, complete urethral coaptation does not occur with pure cellular MDSC therapy, indicating that its effects are at least partially derived from sphincter remodeling and not just a simple bulking effect. Postoperative care should be minimal unless the patient cannot void post-injection, in which case short-term intermittent catheterization should be implemented.

20.6

Current results of clinical studies

Early results of the first clinical studies have recently become available (Table 20.1). Strasser et al. (2007a) reported their comparison of 63 patients undergoing autologous myoblast and fibroblast injection versus 28 patients undergoing collagen injection for SUI. Under ultrasound guidance, a transurethral probe was used to inject (a) fibroblasts into the urethral submucosa to treat mucosal atrophy and (b) myoblasts into the rhabdosphincter for muscle reconstruction (Strasser et al., 2004). The myoblasts and fibroblasts were obtained from an upper arm biopsy. Despite significant increases in Incontinence Quality of Life (I-QOL) instrument scores and decreases in Incontinence Scores in the collagen-treated group, these results did not translate into clinical improvement, and only two female patients (10%) were cured of incontinence. In comparison, 85% of the stem cell-treated group were cured of incontinence, and thickness of urethra and rhabdosphincter were increased significantly at 12 month follow-up on transurethral ultrasonography. Comparing stem cell versus collagen injections in

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Table 20.1 Results from initial clinical studies examining the efficacy of adult stem cell therapy for stress urinary incontinence Number of patients

Indication

Method of delivery

Strasser et al. (2007a)

63

SUI

TUUS

Strasser et al. (2007b)

42

SUI

TUUS

Mitterberger et al. (2007b)

20

SUI

TUUS

Mitterberger et al. (2007a)

123

SUI

TUUS

Mitterberger et al. (2008)

63

PPI

TUUS

Carr et al. (2008)

8

SUI

EI, DPI

Clinical study

Stem cell source

Clinical improvement

Autologous myoblasts and fibroblasts* Autologous myoblasts and fibroblasts* Autologous myoblasts and fibroblasts* Autologous myoblasts and fibroblasts* Autologous myoblasts and fibroblasts* Pure MDSCs

79% – 1 year

91% – 1 year

90% – 1 year 89% – 2 years

79% – 1 year

65% – 1 year

63% – 1 year

TUUS: transurethral ultrasound; EI: endoscopic injection; DPI:, direct periurethral injection; SUI: stress urinary incontinence; PPI: post prostatectomy incontinence. * Mixed cellular plus collagen injections.

female clinical trial participants revealed a 91% incontinence cure rate in the stem cell-treated group compared with 10% in the collagen-treated group (Strasser et al., 2007b). Efficacy at 2 years (89% cure, 11% improvement) was demonstrated in 20 female patients from the same study (Mitterberger et al., 2007). However, in these studies myoblasts and fibroblasts were mixed with 2.5 mL of collagen as carrier material to prevent site migration. The fractional benefit of myoblasts versus fibroblasts versus collagen used in the mixed cellular plus collagen injection approach is unclear. Randomized, controlled clinical trials are necessary to clarify the benefit of the mixed cellular plus collagen injection therapy. Pure cellular clinical therapy with MDSCs obtained from biopsies of the lateral thigh were reported by Carr et al. (2008) and this represented the first trial of North American patients. Eight patients received treatment using either a transurethral or periurethral injection into the middle urethra

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and EUS. The two transurethral injections using a 10 mm needle and the two periurethral injections resulted in measurable improvement, but the two initial injections using the shorter 8 mm needle were not effective. Five of the eight patients with follow-up for over 1 year reported significant improvement occurring between 3 and 8 months after injection (mean follow-up of 16.5 months). In addition, cystoscopy at study exit and surgical exploration in two patients at the time of mid-urethral tape surgery did not demonstrate any appreciable tissue change. Subsequent mid-urethral tape placement and outcome was not negatively impacted by previous MDSCs injection. These results indicate the potential of pure cellular therapy in treating SUI and emphasize the importance of proper cell placement in the effectiveness achieved. Onset of improvement was delayed following injection, suggesting that restoring muscle function may be the mechanism of action in comparison to standard bulking agents. Deeper delivery of MDSCs into the external sphincter appears to be important for successful outcome.

20.7

Conclusions

The transvaginal tape (TVT) procedure has gained popularity for the treatment of SUI. Several authors have reported on the surgical outcome of TVT procedures, demonstrating 85–89% objective cure rates at 3 or 5 years (Ulmsten et al., 1999, Nilsson et al., 2001, Jeffry et al., 2001). In contrast, Ward and Hilton (2004) published the largest randomized, controlled trial, comparing subjective and objective outcomes after abdominal colposuspension or insertion of TVT. At 2 years, 63% of the TVT group and 51% of the colposuspension group were objectively cured; the subjective cure rates at 2 years, however, were only 43% and 37%, respectively. Suburethral sling procedures are regarded as a hammock to reinforce the weakness of pelvic floor muscles and supportive ligaments or fascia, whereas stem cell injection therapy into the middle urethra may restore the contractile response of the striated muscle and rhabdosphincter. The hope is that stem cell treatment of SUI will result in improved cure rates with minimal risks. Not all cellular therapies are the same, as demonstrated by the differences in safety and efficacy among MDSCs, myoblasts, and fibroblasts. Autologous MDSC and ADSC pure injection therapy may be a promising treatment to restore urethral sphincter function.

20.8

Acknowledgements

M.B. Chancellor is a consultant to Cook and has received research grants from Cook. Our SUI research has been supported by National Institutes of Health grants (DK67226, AR49398, and DK55387) and Cook MyoSite funding.

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ward, k. l. and hilton, p. (2004) A prospective multicenter randomized trial of tension-free vaginal tape and colposuspension for primary urodynamic stress incontinence: two-year follow-up. Am J Obstet Gynecol, 190, 324–31. wheeler, e. f., gong, h., grimes, r., benoit, d. and vazquez, l. (1998) p75NTR and Trk receptors are expressed in reciprocal patterns in a wide variety of non-neural tissues during rat embryonic development, indicating independent receptor functions. J Comp Neurol, 391, 407–28. yokoyama, t., pruchnic, r., lee, j. y., chuang, y. c., jumon, h., yoshimura, n., de groat, w. c., huard, j. and chancellor, m. b. (2001a) Autologous primary musclederived cells transfer into the lower urinary tract. Tissue Eng, 7, 395–404. yokoyama, t., yoshimura, n., dhir, r., qu, z., fraser, m. o., kumon, h., de groat, w. c., huard, j. and chancellor, m. b. (2001b) Persistence and survival of autologous muscle derived cells versus bovine collagen as potential treatment of stress urinary incontinence. J Urol, 165, 271–6. zeng, x., jack, g. s., zhang, r. et al. (2006) Treatment of SUI using adipose derived stem cells: Restoration of urethral function. J Urol, 175, 291 [abstract]. zuk, p. a., zhu, m., mizuno, h., huang, j., futrell, j. w., katz, a. j., benhaim, p., lorenz, h. p. and hedrick, m. h. (2001) Multilineage cells from human adipose tissue: implications for cell-based therapies. Tissue Eng, 7, 211–28.

21 Regenerative medicine of the urinary sphincter via direct injection R. Y I O U, CHU Henri Mondor, France

Abstract: Transplantation of cultured muscle precursor cells (MPCs) has recently emerged as a promising therapeutic strategy for stress urinary incontinence aimed at increasing striated urethral sphincter contractions. Satellite cells attached to myofibers of skeletal muscles represent the main source of MPCs capable of reconstituting ad integrum a loss of muscle tissue after an injury. Here, we present the rationale of a new method of intraurethral MPC delivery whereby freshly isolated myofibers and their satellite cells are implanted without prior culture in order to preserve their myogenic potential. After reviewing issues encountered with MPC-based cell therapies in other settings, we describe the results obtained with direct implantation of non-cultured MPCs in the pig. In the pig, direct intraurethral implantation of myofibers with their attached satellite cells leads to a myogenic process consisting of myofiber degeneration rapidly followed by satellite cell activation and myotube formation replacing the parental myofiber. Histological and functional studies showed that the regenerated muscle is innervated by urethral nerves and exerts a tonic activity. Direct delivery of satellite cells without prior cell culture represents an alternative to traditional methods of MPC delivery as a means of engendering muscle formation in the urethra without any cell manipulation. Key words: cell therapy, urinary incontinence, muscle precursor cells, myofiber, sphincter.

21.1

Introduction

Transplantation of muscle precursor cells (MPCs) into the urethra has recently shown potential as a new treatment for stress urinary incontinence (Cannon et al., 2003, Yiou et al., 2003b, Peyromaure et al., 2004). Experimental studies in rodent species have demonstrated significant improvement of striated urethral sphincter contractions after injections of autologous MPCs expanded in culture (Cannon et al., 2003, Yiou et al., 2003b). Recently, Strasser et al., reported promising results in incontinent patients using injections of MPCs and fibroblasts suspended in collagen (Strasser et al., 2007). 445

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The MPCs consist essentially of the satellite cells located under the basal lamina of each myofiber (i.e. the multinucleated contractile cell forming skeletal muscles). After muscle damage, satellite cells proliferate and differentiate into myoblasts that fuse with the pre-existing myofibers or with other myoblasts to produce myotubes (regenerated myofibers). In vitro, live myofibers retain their satellite cells in a quiescent state; their death is a triggering signal for satellite cell proliferation (Bischoff, 1990). Satellite cells confer to skeletal myofibers a remarkable capacity to recover from severe injury. For that reason, injection of such cells has found application in several muscle diseases including genetically determined myopathy (Mendell et al., 1995), heart failure (Menasche et al., 2001) and more recently striated urethral sphincter insufficiency (Strasser et al., 2007). Here, we present the rationale behind a new method of intraurethral MPC delivery consisting of implanting freshly isolated myofibers with their satellite cells without prior expansion in culture in order to avoid cellular alteration. After reviewing issues encountered with MPC-based cell therapies in other settings, we describe the results obtained with direct implantation of noncultured MPCs.

21.2

Challenges with muscle precursor cell transfer

The current method of reference for MPC delivery into a diseased muscle consists of the injection of cells obtained from a muscle biopsy by enzymatic digestion and expanded in vitro. After pioneering work in the 1980s (Partridge et al., 1989), considerable hope was placed in MPC injection as a treatment for several acquired or genetic muscular diseases. Over the following decades numerous papers were published, leading to clinical trials in several conditions such as Duchenne muscular dystrophy (Tremblay et al., 1993, Mendell et al., 1995) or heart failure (Menasche et al., 2001). However, after more than 20 years of research, important issues are still to be addressed to improve this therapeutic approach; the results of injecting a myoblast suspension into diseased muscle have always been so disappointing that this approach is now virtually abandoned for genetic myopathies, the group of conditions for which it was initially designed. Moreover, there is still a lack of clinical evidence that myoblast injection can improve muscle strength. Injecting high doses of myoblasts failed to improve biceps contraction in patients with genetic muscle disease (Mendell et al., 1995). Myoblast transfer used to treat heart failure also failed to improve ventricular contractions; interestingly, the current view is that heart function improvements achieved after myoblast injections are mediated by a paracrine effect and remodeling of the scarred myocardium, rather than by the contractile activity of the injected cells (Menasche, 2007a, Menasche, 2007b).

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The many factors that explain previous failures of myoblast injections include immune reactivity against myoblasts, poor cell survival and limited cell migration (Urish et al., 2005). Using traditional methods of MPC obtention, it was found that less than 3% of the cells are still present 1 hour after injection into a muscle (Beauchamp et al., 1999); actually, only a small subset of MPC retaining stem cell properties has the capacity to survive, explaining why several million cells must be injected to obtain a significant effect. The fact that only myoblasts with stem cell properties survive to injection (Beauchamp et al., 1999) prompted researchers to investigate new preparation methods that select MPCs exhibiting phenotypic characteristics of stem cells. However, clinical studies using such selected cells are still lacking. In parallel, other groups investigated the impact of the cell preparation process on MPC survival. Animal studies comparing injections of MPCs exposed or not exposed to culture conditions are sparse, but their results consistently show adverse effects of culture conditions (Smythe and Grounds, 2000, Collins et al., 2005, Montarras et al., 2005). Smythe and Grounds (2000) found that the initial steps of enzymatic digestion and in vitro expansion were responsible for the rapid death of the MPCs following injection. Subsequently, methods of MPC transplantation bypassing the in vitro steps were investigated with the aim of increasing efficiency. Suzuki et al. (2002) injected four freshly isolated, non-cultured myofibers with attached satellite cells into a murine model of heart damage and found evidence of myotube formation that improved cardiac function. Later, the same group confirmed the high myogenic potential of single myofiber implants by showing that as few as seven satellite cells associated with one transplanted myofiber generate more than 100 new myofibers, provided they do not come into contact with disagregration enzymes (Collins et al., 2005). After enzymatic separation from their parental myofibers, MPCs were found to be several thousand-fold less efficient in generating muscle. Finally, the main challenges of MPC transfer into a diseased muscle are the poor survival of the cells and the loss of myogenic potential associated with the cell preparation process.

21.3

The direct myofiber implantation procedure

Considering these recent concepts, we recently investigated a new method of MPC transfer into the urethra consisting of direct implantation of the myofiber with non-manipulated satellite cells (Lecoeur et al., 2007). The rationale behind this strategy is the fact that the limited number of satellite cells contained in each muscle (2 × 107 to 3 × 107 myonuclei/g of muscle in humans) is capable of reconstituting the totality of the myofibers lost after an injury (Zammit et al., 2002, Morgan and Partridge, 2003). For instance,

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it has been shown that the injection of a myotoxic substance such as notexin into the striated urethral sphincter (Yiou et al., 2003a) or into limb muscles (Sharp et al., 1993) causes rapid and total necrosis of myofibers with concomitant activation of intrinsic satellite cells; full histological and functional recovery, resulting from satellite cell fusion into myotubes, is achieved within 3 weeks. By using muscle implants, we hypothesized that the death of the myofibers following implantation would induce in vivo activation of their satellite cells and myotubes formation. The direct myofiber implantation procedure was investigated in the female pig which was found to have a striated sphincter functionally similar to that in humans (Zini et al., 2006). Each animal underwent a muscle biopsy from which myofibers were isolated and immediately implanted into the proximal third of the urethra (day 0). We previously demonstrated the absence of striated muscle at this location. The striated urethral sphincter is found at the distal third of the urethra, and this allowed the use of antibodies specific for skeletal myosin heavy chains (MHCs) to detect myotubes at the site of injection. Myofibers were first implanted as muscle cores to test the concept and then as muscle strips around the urethra; the latter method aimed to reconstitute a sphincter-like structure. We demonstrated that the myofiber cores act like a reservoir of satellite cells which proliferate immediately after their isolation and fuse in 2–3 weeks to form numerous myotubes replacing the parental myofiber within the urethra. After implantation, myofibers underwent a slow degeneration process that persisted up to day 7. On day 30, the myotubes expressed MHC and were oriented in the same way as the parental myofiber; their myonuclei were in a central position, in contrast to those of the implants immediately after isolation which were located in a peripheral position. The centronucleation process is a common feature of regenerated myofibers (Sharp et al., 1993, Yiou et al., 2003a, Collins et al., 2005). There was no evidence of abscess formation and inflammatory cells were only rarely present at day 30 after implantation. Urodynamic studies showed that implantation of muscle strip around the proximal urethra resulted in a distinct pressure peak in the initial segment of the urethral pressure profilometry. Interestingly, this pressure peak was sensitive to curare injection showing that it was the result of genuine muscular tonic activity under neural control (curare blocks the neuromuscular junctions of striated muscle). This effect is unlikely to be the result of a simple bulking agent, which would not have been modified by curare injection. Myofiber implantation clearly exerted a trophic effect on the urethral nervous system. One month after myofiber implantation, bundles of nerve fibers expressing both neurofilament and VAChT (vesicular acetylcholine transporter) were observed running toward clusters of graft-derived

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(a)

(b)

21.1 (a): Immunostaining for MHC (Alexa Fluor 488) of a cross-section of one pig urethra 1 month after myofiber implantation shows clusters of myotubes arranged in a circular layer. Nuclei are stained with 4,6-diamidino-2-phenylindole (DAP). (b) Staining with bungarotoxin (Alexa-Fluor 488) and immunostaining for neurofilament (Alexa-Fluor 594) on an adjacent cross-section shows acetylcholine receptors connected to newly formed myotubes (arrows shows nerve fibers connected to acetylcholine receptors). Magnification: (a), ×100; (b), ×250.

myotubes (see Fig. 21.1). Nerve density was greater in the vicinity of the myotubes than in other urethral areas deprived of regenerated muscle, where only sparse nerve fibers were observed. The cellular and molecular mechanisms by which regenerating muscle induces nerve sprouting in the urethra probably involve several growth factors produced by myofibers. Insulin-like growth factor, neuroleukin and neural cell adhesion molecule have all been identified as muscle-derived retrograde signaling molecules capable of inducing motorneuron sprouting after partial muscle denervation or direct myofiber damage (van Mier and Lichtman, 1994, English, 2003).

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The regenerated myotubes preferentially adopted a type II phenotype even when the parental myofibers were mainly of type I. These results can be explained by the fact that myofiber phenotype is regulated by the motor innervation. For instance, research on muscle cross-innervation has shown that connection of a slow nerve to a denervated fast muscle induces a shift in the myofiber composition towards phenotype I (Bacou et al., 1996). Our urodynamic and histological findings suggest that the sprouting nerves conferred on the regenerating myotubes a metabolism compatible with sustaining tonic contractions, and that this phenomenon is independent of the phenotype of the parental myofibers. The increase in intraurethral pressure may have been the result of both type I and type II myotubes, specific subtypes of which (IIa) are known to exert a certain tonic action. The concept of cell therapy using direct myofiber implanation differs from other ongoing protocols in which a suspension of mononucleated cells is injected within the incompetent sphincter. The myofiber implantation procedure is original because it involves the implantation of intact myofibers with both satellite cells and surrounding connective tissue. The latter contains fibroblasts, endothelial cells and stem cells involved in the renewal of the satellite cell compartment (Asakura et al., 2002). Interestingly, the regenerated myotubes were oriented longitudinally, in the same direction as the implants. This suggests that satellite cell fusion was tutored by the tubes of the extracellular matrix surrounding the parental myofibers. The orientation of the myotube formation process is a factor that is not controlled when MPCs are injected in a suspension. Finally, the myofiber implantation procedure is close to a transposition of gracilis muscle with preserved neurovascular pedicle (graciloplasty) (Janknegt et al., 1992), the fundamental difference resides in the degeneration/regeneration process associated with myofiber implantation, resulting in the innervation of myotubes by urethral nerves.

21.4

Direct injection of muscle precursor cells using minced muscle

The injection of muscle tissue minced into a fine slurry represents another method of MPC transfer aimed at avoiding in vitro steps of cellular expansion. Formerly, this procedure has been suspected of inducing a strong inflammatory response triggered by myofiber debris that increases MPC death (Pouzet et al., 2000). However, Bierinx and Sebille (2008) used this method in a mouse model of striated sphincter damage. It was found that fragments of minced muscle regenerate new myofibers filling the gap resulting from the trans-section of the striated sphincter. Further studies in large animals are warranted to validate this approach of MPC transfer and rule out side effects possibly associated with myofiber debris in the urethra.

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451

Conclusions and future trends

The myofiber implantation procedure and the injection of minced muscle are original and simple means for transferring MPCs in the urethra. Such approaches of cell therapy may represent alternatives to traditional methods of MPC injection using cultured cells and requiring complex organization in a clinical setting. Further research is required to compare the effects of these different methods of intraurethral MPC delivery. At present a clinical trial is investigating the myofiber implantation procedure for stress urinary incontinence (information available on: http://clinicaltrials.gov/ct2/show/ NCT00472069?term=yiouandrank=1).

21.6

References

asakura, a., seale, p., girgis-gabardo, a. and rudnicki, m. a. (2002) Myogenic specification of side population cells in skeletal muscle. J Cell Biol, 159, 123–34. bacou, f., rouanet, p., barjot, c., janmot, c., vigneron, p. and d’albis, a. (1996) Expression of myosin isoforms in denervated, cross-reinnervated, and electrically stimulated rabbit muscles. Eur J Biochem, 236, 539–47. beauchamp, j. r., morgan, j. e., pagel, c. n. and partridge, t. a. (1999) Dynamics of myoblast transplantation reveal a discrete minority of precursors with stem celllike properties as the myogenic source. J Cell Biol, 144, 1113–22. bierinx, a. s. and sebille, a. (2008) Mouse sectioned muscle regenerates following auto-grafting with muscle fragments: a new muscle precursor cells transfer? Neurosci Lett, 431, 211–14. bischoff, r. (1990) Interaction between satellite cells and skeletal muscle fibers. Development, 109, 943–52. cannon, t. w., lee, j. y., somogyi, g., pruchnic, r., smith, c. p., huard, j. and chancellor, m. b. (2003) Improved sphincter contractility after allogenic musclederived progenitor cell injection into the denervated rat urethra. Urology, 62, 958–63. collins, c. a., olsen, i., zammit, p. s., heslop, l., petrie, a., partridge, t. a. and morgan, j. e. (2005) Stem cell function, self-renewal, and behavioral heterogeneity of cells from the adult muscle satellite cell niche. Cell, 122, 289–301. english, a. w. (2003) Cytokines, growth factors and sprouting at the neuromuscular junction. J Neurocytol, 32, 943–60. janknegt, r. a., baeten, c. g., weil, e. h. and spaans, f. (1992) Electrically stimulated gracilis sphincter for treatment of bladder sphincter incontinence. Lancet, 340, 1129–30. lecoeur, c., swieb, s., zini, l., riviere, c., combrisson, h., gherardi, r., abbou, c. and yiou, r. (2007) Intraurethral transfer of satellite cells by myofiber implants results in the formation of innervated myotubes exerting tonic contractions. J Urol, 178, 332–7. menasche, p. (2007a) [Cellular therapy in cardiology.]. C R Biol, 330, 550–6. menasche, p. (2007b) [Perspectives in cardiac cell therapy]. Presse Med, 36, 1007–11.

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menasche, p., hagege, a. a., scorsin, m., pouzet, b., desnos, m., duboc, d., schwartz, k., vilquin, j. t. and marolleau, j. p. (2001) Myoblast transplantation for heart failure. Lancet, 357, 279–80. mendell, j. r., kissel, j. t., amato, a. a., king, w., signore, l., prior, t. w., sahenk, z., benson, s., mcandrew, p. e., rice, r. et al. (1995) Myoblast transfer in the treatment of Duchenne’s muscular dystrophy. N Engl J Med, 333, 832–8. monterras, d., morgan, j., collins, c., relaix, f., zaffran, s., cumano, a., partridge, t. and buckingham, m. (2005) Direct isolation of satellite cells for skeletal muscle regeneration. Science, 309, 2064–7. morgan, j. e. and partridge, t. a. (2003) Muscle satellite cells. Int J Biochem Cell Biol, 35, 1151–6. partridge, t. a., morgan, j. e., coulton, g. r., hoffman, e. p. and kunkel, l. m. (1989) Conversion of mdx myofibres from dystrophin-negative to -positive by injection of normal myoblasts. Nature, 337, 176–9. peyromaure, m., sebe, p., praud, c., derocle, g., potin, n., pinset, c. and sebille, a. (2004) Fate of implanted syngenic muscle precursor cells in striated urethral sphincter of female rats: perspectives for treatment of urinary incontinence. Urology, 64, 1037–41. pouzet, b., vilquin, j. t., hagege, a. a., scorsin, m., messas, e., fiszman, m., schwartz, k. and menasche, p. (2000) Intramyocardial transplantation of autologous myoblasts: can tissue processing be optimized? Circulation, 102, III 210–15. sharp, n. j., kornegay, j. n., bartlett, r. j., hung, w. y. and dykstra, m. j. (1993) Notexin-induced muscle injury in the dog. J Neurol Sci, 116, 73–81. smythe, g. m. and grounds, m. d. (2000) Exposure to tissue culture conditions can adversely affect myoblast behavior in vivo in whole muscle grafts: implications for myoblast transfer therapy. Cell Transplant, 9, 379–93. strasser, h., marksteiner, r., margreiter, e., pinggera, g. m., mitterberger, m., frauscher, f., ulmer, h., fussenegger, m., kofler, k. and bartsch, g. (2007) Autologous myoblasts and fibroblasts versus collagen for treatment of stress urinary incontinence in women: a randomised controlled trial. Lancet, 369, 2179–86. suzuki, k., murtuza, b., heslop, l., morgan, j. e., smolenski, r. t., suzuki, n., partridge, t. a. and yacoub, m. h. (2002) Single fibers of skeletal muscle as a novel graft for cell transplantation to the heart. J Thorac Cardiovasc Surg, 123, 984–92. tremblay, j. p., malouin, f., roy, r., huard, j., bouchard, j. p., satoh, a. and richards, c. l. (1993) Results of a triple blind clinical study of myoblast transplantations without immunosuppressive treatment in young boys with Duchenne muscular dystrophy. Cell Transplant, 2, 99–112. urish, k., kanda, y. and huard, j. (2005) Initial failure in myoblast transplantation therapy has led the way toward the isolation of muscle stem cells: potential for tissue regeneration. Curr Top Dev Biol, 68, 263–80. van mier, p. and lichtman, j. w. (1994) Regenerating muscle fibers induce directional sprouting from nearby nerve terminals: studies in living mice. J Neurosci, 14, 5672–86. yiou, r., lefaucheur, j. p. and atala, a. (2003a) The regeneration process of the striated urethral sphincter involves activation of intrinsic satellite cells. Anat Embryol (Berl), 206, 429–35. yiou, r., yoo, j. j. and atala, a. (2003b) Restoration of functional motor units in a rat model of sphincter injury by muscle precursor cell autografts. Transplantation, 76, 1053–60.

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zammit, p. s., heslop, l., hudon, v., rosenblatt, j. d., tajbakhsh, s., buckingham, m. e., beauchamp, j. r. and partridge, t. a. (2002) Kinetics of myoblast proliferation show that resident satellite cells are competent to fully regenerate skeletal muscle fibers. Exp Cell Res, 281, 39–49. zini, l., lecoeur, c., swieb, s., combrisson, h., delmas, v., gherardi, r., abbou, c., chopin, d. and yiou, r. (2006) The striated urethral sphincter of the pig shows morphological and functional characteristics essential for the evaluation of treatments for sphincter insufficiency. J Urol, 176, 2729–35.

22 Regenerative medicine for the urethra T. A B O U S H WA R E B and A. ATA L A, Wake Forest Institute for Regenerative Medicine, USA; and A. E L K A S S A B Y, Ain Shams University, Egypt

Abstract: Replacement of damaged or diseased tissues represents one of the most challenging aspects of medicine. Within the urinary tract, tissue replacement is a technically demanding procedure. In particular, urethral reconstruction has posed a continuing challenge to the reconstructive urologist. Various surgical techniques utilizing flaps or grafts (genital/ extragenital), whether single or multistage, have been used for urethral reconstruction with satisfactory results. Genital skin, however, is not always readily available and most graft harvesting procedures are technically demanding and time consuming. The development of an ‘off-the-shelf’ graft material may be the answer to this old urological dilemma. In this chapter, we review the current materials, whether synthetic or natural, that have been used in clinical or preclinical studies for the purpose of urethral reconstruction. Key words: urethral reconstruction, scaffolds, naturally derived materials, off the-shelf grafts.

22.1

Introduction

Replacement of damaged or diseased tissues represents one of the most challenging aspects of medicine. Within the urinary tract, tissue replacement is a technically demanding procedure. In particular, urethral reconstruction has posed a continuing challenge to the reconstructive urologist. Various surgical techniques utilizing flaps or grafts (genital/extragenital), whether single or multistage, have been used for urethral reconstruction with satisfactory results. The consensus in the urological community has always been that flaps, even though they are more demanding surgically, are superior to grafts in terms of outcome. However, in complex and recurrent cases, an adequate amount of non-hirsute genital skin is not always available. Different types of grafts have been employed – including split-thickness and full-thickness skin, bladder mucosa, and buccal mucosa grafts (which have become the favored graft for urethral substitution) (Bhargava et al., 2004) – but these graft types may not always be sufficient to bridge long urethral defects. 454

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Development of ‘off-the-shelf’ graft materials may solve some of the problems associated with current urological reconstructive techniques. Use of such materials could lead to elimination of donor site morbidity (Tolstunov et al., 1997), shorter operative duration, and decreased blood loss during surgery. In recent years, tissue engineering techniques have been employed to develop this type of ready-to-use graft tissue, and preliminary data indicate that these engineered graft materials may be used to repair urethral defects of unrestricted dimensions, because they assist the patient’s own regenerative processes in recreating a normal urethral wall. More advanced techniques of regenerative medicine are now being developed that allow reconstitution of entire tubularized urethral segments in vitro using a patient’s own cells followed by implantation of these segments for complete urethral substitution. Implantable graft materials, termed matrices or scaffolds, direct new tissue growth in predetermined patterns through modulation of the body’s inherent ability to regenerate (Atala, 2003). In addition, scaffolds provide structural support upon which cells can grow and become vascularized (Atala, 2001, Han and Liu, 1999). Effective repair and regeneration of injured tissues depends on early re-establishment of the blood flow needed for cellular oxygenation and metabolic requirements. Ideally, scaffolds serve as an artificial extracellular matrix (ECM) in which cell behavior can be regulated. The ultimate biomaterial should therefore promote cell adhesion, proliferation, migration, and differentiation, and it should be biocompatible and biodegradable. Finally, it should have appropriate mechanical and physical properties (Kimuli et al., 2004). In addition, some types of scaffolds can facilitate the delivery of bioactive factors (e.g. cell adhesion peptides, growth factors, etc.) to desired sites in the body, which can enhance healing and regeneration of some tissues (Kim and Mooney, 1998). Scaffolds can be synthetic polymers, or they can be derived from natural materials. Natural scaffold materials include acellular tissue matrices (e.g. bladder submucosa and small intestinal submucosa) (Atala, 2003). Commonly used synthetic scaffolds consist of polymers of lactic and glycolic acid. Both types of scaffolding materials have been used in urethral reconstruction to date, whether in experimental or clinical settings. Some of these materials have shown great potential for success. In this chapter, we review the uses of various scaffolds for urological applications, with special focus on our own experience in the field.

22.2

Synthetic scaffolds

Synthetically derived scaffolds are strong, can be fabricated to degrade at predetermined rates following implantation, and can be designed to mimic the material properties of the native tissue they are designed to replace. In

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addition, they can be manufactured consistently on a large scale. While these characteristics are appealing, clinical complications are often encountered when synthetic scaffold materials are used. Many synthetic biopolymers, those based on a natural material such as hyaluronic acid-based polymers, support the growth of cells in dedifferentiated or incompletely redifferentiated states (Aigner et al., 1998) or, like polyglycolic acid (PGA), lead to tissue deposition that is less than optimal (Cao et al., 1998). Synthetic scaffolds made of permanent synthetic biopolymers, e.g. silicone and polytetrafluoroethylene (PTFE), carry the risk of infection, calcification, and the formation of excessive scar tissue (Atala, 1997, Mendelsohn and Dunlop, 1998, Sclafani et al., 1997). As a clinical example, many synthetic biopolymers have been used to replace damaged vascular structures, and while long-term patency rates have risen over the years, the ideal vascular graft scaffold remains elusive. For example, no synthetic biopolymer currently available for clinical use can restore normal structure and function to injured vascular tissues while avoiding severe complications such as thrombosis, neointimal hyperplasia, accelerated atherosclerosis, and/or calcification (Chen et al., 1997, Vyavahare et al., 1998). To date, there has not been a single synthetic material that has been identified that will support the epithelial and smooth muscle cells required to regenerate a functional urethra, and their use has been associated with poor results so far (Atala, 1999). Thus, urethral tissue engineering strategies have focused on the development of biological substitutes.

22.3

Biological (natural) polymers

Another approach to the repair and regeneration of damaged tissues is the use of intact ECM obtained from animal tissues to act as the growth support for host cells. Native biological tissue is composed of two basic components: cells and ECM. ECM refers to the non-cellular part of a tissue that provides support to cells, controls the structure of the tissue, and regulates the cell phenotype (Alberts et al., 1994). ECM consists of protein and carbohydrate structures secreted by the resident cells. The most common constituent of the ECM is the structural protein collagen. When harvested from a tissue source, collagen ECM can be fabricated into graft prostheses by removing the cells through gentle mechanical and chemical processing. The resulting ECM material may be referred to as naturally occurring polymeric scaffolds, bioscaffolds, or biomatrices. These materials, though harvested from several different body systems, all share several characteristics once they are processed into a graft material. Specifically, since they are subjected to minimal processing after they are removed from the source animal, they retain a structure and composition nearly identical to their native state.

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Once the host cells are removed, the scaffolds can be implanted acellularly to replace diseased or damaged tissues. These naturally occurring materials offer promising alternatives to synthetically engineered polymeric scaffolds. Naturally occurring scaffolds can be processed so that growth factors such as basic fibroblast growth factor (FGF-2) and transforming growth factor-β (TGF-β) (Koizumi et al., 2000, Voytik-Harbin et al., 1997), glycosaminoglycans such as heparin and dermatan sulfate (Hodde, 1996, Meinert et al., 2001, Peel et al., 1998), and structural elements such as fibronectin (McPherson and Badylak, 1998), elastin, and collagen are preserved (Chaplin et al., 1999). These materials prevent many of the complications associated with foreign material implants because they provide a natural environment in which cells can attach, migrate, proliferate, and differentiate. These naturally occurring biopolymers have been shown to interact quickly with host tissues. They may induce the deposition of cells and additional ECM, as well as promote rapid angiogenesis. These functions are essential for the restoration of functional soft tissue. Naturally occurring biopolymers include acellular dermal matrix (AlloDerm®), cadaveric fascia, amniotic membrane, small intestinal submucosa (SIS), and bladder acellular matrix graft (BAMG).

22.3.1 Acellular dermal matrix Skin is comprised of two primary layers that differ in function, thickness, and strength: the epidermis and the dermis. The dermis underlies the epidermis and possesses an extensive vascular system that supplies the epidermis with nutrient-rich blood. Additionally, the dermis helps the skin regulate temperature. The dermis is composed of fibroblasts, which produce a collagen-containing ECM that provides elasticity and support to the skin. Acellular dermal matrices for tissue engineering purposes are harvested from either pig skin or human cadaver skin. The epidermis is removed by soaking the skin in 1 mol/l NaCl for 8 hours. Dermal fibroblasts and epithelial cells are removed by incubation of the material in 2% deoxycholic acid containing 10 mmol/l ethylenediaminetetraacetate (EDTA). The dermis is then cryoprotected with a combination of 35% maltodextrin and 10 mmol/l disodium-EDTA, and freeze dried until use (Chaplin et al., 1999). When it has been implanted as an acellular tissue graft, acellular dermis has supported endothelialization of repaired vascular structures (Inoue et al., 1996), inhibited excessive wound contraction (Walden et al., 2000), and supported host cell incorporation and capillary ingrowth into the grafted site (Dalla Vecchia et al., 1999, Medalie et al., 1996). By using acellular de-epidermized dermis (DED), Bhargava et al. (2004) reported the successful culture of tissue-engineered buccal mucosa (TEBM) that closely resembled the native oral mucosa and was suitable for use in substitution urethroplasty.

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22.3.2 Cadaveric fascia lata The tensor fascia lata is the thick band of connective tissue attaching the pelvis to the knee on the lateral side of the leg. At the hip, its muscular components join to thick connective tissues and help stabilize the hip and knee joints by putting tension on the iliotibial band (IT band). The IT band, and the distal section of the tensor fascia lata, is harvested for the graft prosthesis. In its native state, the fascia lata tendon is composed of heavy, parallel bundles of type I collagenous fibers that are held together by ECM tissue. Between the bundles of fibers are fibroblasts, nerves, and blood vessels that supply the tendon with nutrients. Processing of cadaveric fascia involves ethanol extraction followed by high-pressure washing with antibiotics. The extracted tissue is then lyophilized and terminally sterilized with gamma irradiation. Evaluation of the graft material revealed an almost pure collagen tissue structure (Blander and Zimmern, 2000). Intraoperatively, the grafts are reconstituted in a 0.9% saline soak for 3 minutes prior to implantation (Carbone et al., 2001). One common use of fascia lata is in the repair of female stress urinary incontinence and genitourinary prolapse. Although not completely acellular, the relative paucity of cells in this type of graft is probably the major characteristic that allows the successful use of donor fascia without the need for histocompatibility between donor and recipient (Fitzgerald et al., 2000). Furthermore, one study using fresh-frozen donor fascia lata allografts in rabbit cornea showed that after 1 week, donor fascia lata fibroblasts were absent. Within 3 weeks, the grafts were repopulated by host fibroblasts (Crawford, 1969).

22.3.3 Amniotic membrane The amniotic membrane forms the sac that encloses the embryo during pregnancy. It is an extremely strong, 5–20 μm thick tissue that has been used as a graft in several repair applications. In its native state, the epithelium of the amnion consists of a single layer of cells resting upon a relatively cellfree basement membrane (Aplin et al., 1985). This membrane is comprised of several collagen types, including the fibrillar collagen types I and III, and the basal lamina collagen type IV (Aplin et al., 1985, Lei et al., 1999). Several growth factors – including epidermal growth factor, several transforming growth factor isoforms, basic fibroblast growth factor, keratinocyte growth factor, and hepatocyte growth factor – have also been identified in the membrane and it has been reported that these factors are retained in the processed tissue matrix (Koizumi et al., 2000). Amniotic membrane is obtained at parturition and cleaned of blood with saline containing penicillin, streptomycin, amphotericin B, and clindamycin (Avila et al., 2001). It is separated from the chorion by blunt dissection, washed in sterile water, and

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treated by soaking for 3 hours in a 10% solution of trypsin to lyse the cells. The membrane is then sterilized with gamma irradiation and frozen until clinical use (Meezan et al., 1975, Young et al., 1991).

22.3.4 Small intestinal submucosa SIS is an acellular matrix that consists of three distinct layers of the mammalian small intestine: lamina propria, muscularis mucosa, and tunica submucosa (Badylak, 1993). The tunica submucosa is the layer of connective tissue arranged immediately under the mucosa layer of the intestine and is a 100–200 μm thick interstitial ECM; it makes up the bulk of the SIS biopolymer scaffold. In the living intestine, the submucosa supports the mucosal structures and is secreted and maintained by the fibroblasts that reside there (Berne and Levy, 1996). SIS is composed primarily of fibrillar collagens and adhesive glycoproteins (McPherson and Badylak, 1998), which serve as a scaffold into which cells can migrate and multiply. The SIS layers also contain potent regulatory factors – such as glycosaminoglycans, proteoglycans, and growth factors – which regulate cellular processes that maintain tissue homeostasis and respond to injury and infection (Berne and Levy, 1996). Many of these components are retained in the biopolymer scaffold following processing. SIS is harvested from the mammalian small intestine by mechanically separating it from its outer muscular layers and internal mucosal layers. The biopolymer is thoroughly rinsed in water, treated with an aqueous solution of 0.1% peracetic acid, and rinsed in sequential exchanges of water and phosphate-buffered saline to yield a neutral pH (Badylak et al., 2000). It is then either stored in an antibiotic solution containing 0.05% Gentamycin sulfate (Clarke et al., 1996), or sterilized using 2.5 mRad gamma irradiation (Badylak et al., 2000). When SIS is implanted as a naturally occurring biopolymer scaffold, it stimulates angiogenesis as well as connective and epithelial tissue growth and differentiation. It also supports deposition, organization, and maturation of ECM components that are functionally and histologically appropriate for the site of implantation (Badylak, 1993). SIS has also been used experimentally in animals for urethral augmentation (El-Assmy et al., 2004, Kropp et al., 1998, Mantovani et al., 2002, Sievert et al., 2001). Kropp et al. (1998) used SIS as an onlay patch for urethroplasty in rabbits. In this experiment SIS promoted regeneration of three to four layers of stratified columnar epithelium as well as suburothelial circular smooth muscles. El-Assmy et al. (2004) evaluated the use of single-layer SIS for urethral replacement in a rabbit model, as both an onlay and a tubular graft. When SIS was used as a tubular graft, all rabbits developed urethral strictures with progressive narrowing. Six out of nine rabbits that received onlay patch grafts had normal urethrograms, while there was

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relative narrowing in the other three. Histological examination showed epithelial regeneration with no regeneration of smooth muscle in the region of the SIS onlay graft at up to 3 months; instead, fibrous connective tissue was present. Healing was inferior to spontaneously healed, non-patched urethras, which showed regeneration of bundles of smooth muscle circumferentially surrounding the urethral mucosa. In humans, Mantovani et al. (2003) published results from a series of five patients in whom SIS was used for urethral stricture repair. These patients had no recurrences at 6 months after surgery. However, Hauser et al. (2006) used SIS as a dorsal onlay patch urethroplasty in five patients and, in this study, recurrent urethral stricture developed in four patients after a median of 14 months (Weiser et al., 2003).

22.3.5 Bladder acellular matrix graft BAMG is a collagen-based matrix, derived from allogenic bladder submucosa (Merguerian et al., 2000). Unlike the intestinal submucosa, however, which is easily separated from the external muscle layers, the submucosa of the urinary bladder is intimately attached to the muscular bladder wall. Complete mechanical separation of the layers has proven tedious and difficult, and so attempts at rendering the bladder submucosa muscle free have often resorted to chemical and/or enzymatic agents such as sodium hydroxide, sodium desoxycholate, sodium dodecyl sulfate, or deoxyribonuclease (Badylak et al., 2000, Kropp, 1999, Reddy et al., 2000, Wefer et al., 2001). The evaluation of the BAMG for tissue repair has largely been limited to urological applications. In the urinary system, most investigation has focused on the use of BAMG for augmentation cystoplasty and urethral reconstruction. Yoo et al. (1998) showed that bladders replaced with acellular bladder matrices had approximately a 30% increase in capacity, compared with bladders reconstructed with cell-seeded bladder submucosa. When implanted as a replacement for bladder tissue, BAMG has been shown to act as a stimulus for whole-bladder regeneration and restoration of bladder function (Merguerian et al., 2000). It incited the restoration of a urothelium complete with newly formed ECM, underlying smooth muscle cells, and immature blood vessels indicative of a neovascular response (Wefer et al., 2001). Probst et al. (2000) used BAMG for bladder augmentation in dogs and confirmed that the reconstructed bladder matched the morphological and functional properties of the normal bladder Piechota et al. (1998b). In addition, in a rat model, there was regeneration of all cellular components of the bladder (Piechota et al., 1998a, Sutherland et al., 1996). However, Sutherland et al. (1996) reported that the muscle layer, although present, was not fully developed in cases where acellular matrix material was used for reconstruction.

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Oberpenning et al. (1999) compared the outcome of augmentation cystoplasty using acellular matrices or cell-seeded matrices in an animal model after subtotal cystectomy. The cell-seeded matrices were seeded with autologous urothelial and smooth muscle cells. The bladders augmented with acellular matrices maintained average capacities of 46% of the initial values; however, when cell-seeded matrices were used, the reconstructed bladders maintained 95% of the initial capacity. These findings were confirmed radiographically. In addition, a marked decrease in bladder compliance (42%) was measured in the acellular matrix group, while normal compliance was seen in animals from the cell-seeded matrix group. Histologically, the polymer-only bladder tissue consisted of normal urothelial cells with a thickened fibrotic submucosa and a thin layer of muscle fibers. The cellseeded matrices, on the other hand, demonstrated a normal cellular organization, consisting of a triple layer of urothelium, submucosa, and muscle. In other experiments, Obara et al. (2006) assessed the feasibility of bladder acellular matrix (BAM) grafting on to the bladder of rats with spinal cord injury (SCI). Regenerated urothelium, smooth muscles, and nerve fibers in the grafted BAM appeared at 2, 4, and 12 weeks, respectively, in both intact and SCI rats. Immunohistological examination showed that these regenerated tissues inherited each characteristic of the host bladder tissue and the grafted BAM showed proper distention to store urine at 12 weeks, suggesting that this particular technique could be used for patients with neurogenic bladder who also require augmentation enterocystoplasty. Tissue-engineered bladders have also been used clinically. Atala et al. (2006) employed tissue-engineered autologous bladders for cystoplasty in seven patients with myelomeningocele and end-stage bladder disease. Autologous urothelial and smooth muscle cells were seeded on a biodegradable bladder-shaped scaffold made of collagen, or a composite of collagen and PGA. The autologous-engineered bladder constructs were then implanted either with or without an omental wrap. After a mean follow-up of 46 months, the mean volume and compliance increased 2.8-fold in these patients. Various bladder acellular scaffolds have also been used experimentally for the regeneration of urethral tissue (Atala et al., 1999, Chen et al., 1999, Sievert et al., 2000). In some animal studies, segments of the urethra were resected and replaced with BAMG in an onlay fashion. Histological examination showed complete epithelialization and progressive vessel and muscle infiltration, and the animals were able to void through the neourethras with no signs of strictures or complications (Chen et al., 1999). Sievert et al. (2000) showed that, in a rabbit model, all tissue components were seen in the grafted matrix after 3 months. However, the amount of smooth muscle tissue in the matrix was less than that seen in the normal rabbit urethra.

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(c) (d)

22.1 Tissue engineering of the urethra using a collagen matrix. (a) Representative case of a patient with a bulbar stricture. (b) During the urethral repair surgery, strictured tissue is excised, preserving the urethral plate on the left side, and matrix is anastamosed to the urethral plate in an onlay fashion on the right. (c) Urethrogram 6 months after repair. (d) Cystoscopic view of urethra before surgery on the left side, and 4 months after repair on the right side.

In humans, our center reported on the uses of homologous BAMG in the treatment of urethral pathology (Atala et al., 1999). We used the same techniques for decellularization of the bladder submucosa as we did for the bladder reconstruction studies (Yoo et al., 1998). The acellular bladder submucosa was applied in a series of patients with a history of failed hypospadias reconstruction, in which the urethral defects were repaired with an onlay patch (Fig. 22.1)(Atala et al., 1999). Three out of the four patients had successful outcomes. Following this, we applied the same material clinically in the treatment of urethral stricture disease (El-Kassaby et al., 2003). Stricture length ranged from 1.5 to 16 cm. The results were encouraging, as 24 of the 28 patients had a successful outcome after a mean follow-up of 37 months. It is possible that this high success rate could be a result of the careful selection of patients for study, as this was not a randomized trial. In a third study, we employed BAMG as an onlay patch in five children with urethral stricture and insufficient skin for traditional grafting procedures. Successful outcome was reported in four of these children after 30 months (El-Kassaby et al., 2002). More recently, we conducted a fourth clinical study using BAMG. A prospective randomized comparative study between buccal

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mucosa and BAMG grafts for the management of urethral stricture disease was completed (El-Kassaby et al., 2008). The study included 30 patients with anterior urethral strictures and lack of native autologous tissue available for repair. Patients were randomly divided between those who would receive buccal mucosa grafts and those who would receive BAMG repairs. The graft was obtained either off-the-shelf (BAMG) or from the oral cavity through the accepted harvest technique for buccal mucosa. The graft was then tailored to the required length and width of the stricture and was applied on the floor of the urethra in an onlay fashion. Surgical outcome was successful in all buccal mucosal patients with a success rate of 100%. In the BAMG group of patients, success rate varied according to the number of previous surgical operations. From the nine patients with fewer than two previous operations, eight had a successful outcome and the ninth required an end-to-end anastomosis procedure for a recurrent anastomotic stricture. In the other group of BAMG (with two or more previous interventions), only two patients (33%) out of the six had a successful outcome while the other four required interventions postoperatively. Follow-up of all patients ranged from 18 to 36 months (mean 25 months). In the buccal mucosal group, the nature of the urethral edges did not seem to have an impact on the outcome of the procedure and all patients had successful outcomes. In the BAMG group, however, the nature of the urethral edges had a definite role in determining the outcome. The spongiofibrosis and the lack of healthy edges adversely affected the mucosal growth over the scaffold, which depends totally on the quality of the surrounding urethral mucosa. Therefore, it is evident that the acellular bladder submucosa can be used as a dorsal onlay graft in urethral reconstruction if there is a healthy urethral bed present that would allow cell migration from the periphery. The above findings were confirmed by a series of other experiments. These acellular collagen matrices were applied as an onlay patch on the ‘ventral’ aspect of the urethra. This suggests that the use of acellular scaffold grafts in ‘dorsal’ onlay patch procedures, as reported by Hauser et al. (2006), is neither necessary nor helpful in urethral stricture disease because the edges of the dorsally incised urethra can be fixed to the tunica albuginea, which covers the underside of the corpora, and the urethral mucosa will regenerate in that position with or without an acellular scaffold (Monseur, 1968). Interestingly, Shokeir et al. (2003) reported the use of urethral acellular matrix (organ-specific) applied as a ventral onlay patch in a dog model. They compared regeneration over the acellular matrix (study group) to a control group in which no matrix was implanted. Histopathological examination showed gradual regeneration over the acellular matrix with normal appearance at 20 weeks. In the control group, normal healing was observed at 2 months. They concluded that regeneration of all components of the urethra could occur with or without acellular matrix. This is attributed to

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the fact that the experiment was performed using normal urethral tissue that is capable of spontaneous regeneration. In essence, this follows the principles of Russell’s technique (Russell, 1914) and the Denis–Browne buried skin strip procedure, which require a healthy tissue strip to regenerate into a tube over a splinting catheter. This highlights the importance of verifying the presence of healthy urethral edges that are capable of regeneration after stricturotomy before adopting an acellular graft onlay patch, so as to optimize the subsequent regenerative process. If the strip is unhealthy, the outcome will be poor. In these patients, excision of most of the fibrotic segment and roof strip anastomosis of the urethra is performed according to the Russell technique (Russell, 1914). This decreases the length of the patch graft required and, more importantly, will guarantee the presence of healthy urethral edges that are capable of regeneration. Our group also showed that although acellular collagen-based grafts may be suitable for partial urethral repair in an onlay fashion, they are ineffective for replacing tubularized segments. De Filippo et al. (2002) suggested the use of tubularized collagen matrices, seeded with autologous urothelial and muscle cells to avoid stricture in the neourethras. The use of autologous cells involves obtaining a biopsy of tissue from the host, dissociating the cells from the biopsy tissue, and expanding them in vitro. Cells are then reattached to a matrix, and reimplanted into the same host (De Filippo et al., 2002, Olsen et al., 1992). These seeded collagen matrices have also been used for total penile urethral replacement (De Filippo et al., 2002). Recently, the Atala group (Orabi et al., 2007) (unpublished data, 2007) have conducted a larger animal study for complete tubularized urethral stricture repair in dogs. The group tested the technique in a longer segment (6–7 cm) of the canine urethra. Both seeded and unseeded matrix replacements were performed, and the animals were studied at 1, 3, and 6 month time points. The animals were examined with three-dimensional computed tomography scans over the course of the study, and ascending and micturating urethrocystograms were performed. The study indicated that seeded BAMG can maintain a normal caliber urethra after complete tubularized replacement for up to 6 months in the dog model. The acellular matrix replacements all failed, indicating that cell seeding is essential in maintaining a viable urethral segment in tubularized grafts (Orabi et al., 2007).

22.4

Conclusions

The reconstructive strategies for patients with urethral disease using regenerative medicine techniques depend on the status of the urethra. Patients who can be reconstructed with an onlay type of repair may benefit from acellular bladder submucosa grafts as long as the urethral bed is healthy, allowing an adequate anchor for cell migration. If the urethral bed is

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unhealthy a dorsal onlay graft with cellular tissues – such as bladder submucosa, skin, or cell-seeded grafts – may be preferable. If a tubularized repair is required, cell seeded grafts are needed. The success of using acellular matrix scaffold grafts for urethral reconstruction depends on proper selection of host tissue so that conditions for regeneration of all urethral layers are maximized. Because the urethra functions as a conduit, its structure makes it amenable to replacement by tissue engineering and may make the off-the-shelf graft a real option for the urologist.

22.5

Acknowledgement

The authors wish to thank Dr Jennifer Olson for editorial assistance with this chapter.

22.6

References

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oberpenning, f., meng, j., yoo, j. j. and atala, a. (1999) De novo reconstitution of a functional mammalian urinary bladder by tissue engineering. Nat Biotechnol, 17, 149–55. olsen, l., bowald, s., busch, c., carlsten, j. and eriksson, i. (1992) Urethral reconstruction with a new synthetic absorbable device. An experimental study. Scand J Urol Nephrol, 26, 323–6. orabi, h. a., aboushwareb, t. a., zhang, y., yoo, j. j. and atala, a. (2007) Tissue engineered tubularized urethra for surgical reconstruction: a pre-clinical study. AUA Annual Meeting, Anaheim, California. peel, s. a. f., chen, h., renlund, r., badylak, s. f. and kandel, r. a. (1998) Formation of a SIS-cartilage composite graft in vitro and its use in the repair of articular cartilage defects. Tissue Eng, 4, 143. piechota, h. j., dahms, s. e., nunes, l. s., dahiya, r., lue, t. f. and tanagho, e. a. (1998a) In vitro functional properties of the rat bladder regenerated by the bladder acellular matrix graft. J Urol, 159, 1717–24. piechota, h. j., dahms, s. e., probst, m., gleason, c. a., nunes, l. s., dahiya, r., lue, t. f. and tanagho, e. a. (1998b) Functional rat bladder regeneration through xenotransplantation of the bladder acellular matrix graft. Br J Urol, 81, 548–59. probst, m., piechota, h. j., dahiya, r. and tanagho, e. a. (2000) Homologous bladder augmentation in dog with the bladder acellular matrix graft. BJU Int, 85, 362–71. reddy, p. p., barrieras, d. j., wilson, g., bagli, d. j., mclorie, g. a., khoury, a. e. and merguerian, p. a. (2000) Regeneration of functional bladder substitutes using large segment acellular matrix allografts in a porcine model. J Urol, 164, 936–41. russell, r. h. (1914) The treatment of urethral stricture by excision. Br J Surg, 2, 375. sclafani, a. p., thomas, j. r., cox, a. j. and cooper, m. h. (1997) Clinical and histologic response of subcutaneous expanded polytetrafluoroethylene (Gore-Tex) and porous high-density polyethylene (Medpor) implants to acute and early infection. Arch Otolaryngol Head Neck Surg, 123, 328–36. shokeir, a., osman, y., el-sherbiny, m., gabr, m., mohsen, t. and el-baz, m. (2003) Comparison of partial urethral replacement with acellular matrix versus spontaneous urethral regeneration in a canine model. Eur Urol, 44, 603–9. sievert, k. d., bakircioglu, m. e., nunes, l., tu, r., dahiya, r. and tanagho, e. a. (2000) Homologous acellular matrix graft for urethral reconstruction in the rabbit: histological and functional evaluation. J Urol, 163, 1958–65. sievert, k. d., wefer, j., bakircioglu, m. e., nunes, l., dahiya, r. and tanagho, e. a. (2001) Heterologous acellular matrix graft for reconstruction of the rabbit urethra: histological and functional evaluation. J Urol, 165, 2096–102. sutherland, r. s., baskin, l. s., hayward, s. w. and cunha, g. r. (1996) Regeneration of bladder urothelium, smooth muscle, blood vessels and nerves into an acellular tissue matrix. J Urol, 156, 571–7. tolstunov, l., pogrel, m. a. and mcaninch, j. w. (1997) Intraoral morbidity following free buccal mucosal graft harvesting for urethroplasty. Oral Surg Oral Med Oral Pathol Oral Radiol Endod, 84, 480–2. voytik-harbin, s. l., brightman, a. o., kraine, m. r., waisner, b. and badylak, s. f. (1997) Identification of extractable growth factors from small intestinal submucosa. J Cell Biochem, 67, 478–91.

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vyavahare, n. r., hirsch, d., lerner, e., baskin, j. z., zand, r., schoen, f. j. and levy, r. j. (1998) Prevention of calcification of glutaraldehyde-crosslinked porcine aortic cusps by ethanol preincubation: mechanistic studies of protein structure and water-biomaterial relationships. J Biomed Mater Res, 40, 577–85. walden, j. l., garcia, h., hawkins, h., crouchet, j. r., traber, l. and gore, d. c. (2000) Both dermal matrix and epidermis contribute to an inhibition of wound contraction. Ann Plast Surg, 45, 162–6. wefer, j., sievert, k. d., schlote, n., wefer, a. e., nunes, l., dahiya, r., gleason, c. a. and tanagho, e. a. (2001) Time dependent smooth muscle regeneration and maturation in a bladder acellular matrix graft: histological studies and in vivo functional evaluation. J Urol, 165, 1755–9. weiser, a. c., franco, i., herz, d. b., silver, r. i. and reda, e. f. (2003) Single layered small intestinal submucosa in the repair of severe chordee and complicated hypospadias. J Urol, 170, 1593–5; disussion 1595. yoo, j. j., meng, j., oberpenning, f. and atala, a. (1998) Bladder augmentation using allogenic bladder submucosa seeded with cells. Urology, 51, 221–5. young, r. l., cota, j., zund, g., mason, b. a. and wheeler, j. m. (1991) The use of an amniotic membrane graft to prevent postoperative adhesions. Fertil Steril, 55, 624–8.

23 Penile reconstruction H.- J. WA N G and J. J. YO O, Wake Forest Institute for Regenerative Medicine, USA

Abstract: Conditions such as congenital anomalies of genitalia, penile cancer, traumatic penile injury and some vasculogenic erectile dysfunctions often require extensive surgical procedures to correct anatomical and functional deformities. The ultimate goal of these procedures is to restore a cosmetically acceptable penis that allows normal urinary, sexual and reproductive function. However, reconstruction of the penis has consistently been a challenge, primarily owing to the shortage of native tissue and inability to restore normal tissue function. In recent years, tissue engineering techniques have been developed to generate biological substitutes that could be used to restore and maintain normal tissue function. This chapter presents various tissue engineering and regenerative medicine approaches that aim to achieve functional penile tissue. Key words: penile, tissue engineering, autologous cells, reconstructive surgery, stem cells.

23.1

Introduction

Various penile pathologies often require reconstructive procedures in order to achieve aesthetic and functional restoration. These pathologies encompass all age groups with conditions ranging from congenital anomaly to acquired disease, such as carcinoma, trauma, Peyronie’s disease and severe erectile dysfunction (Ralph et al., 2006). Reconstruction of the penis has consistently been a challenge, primarily because the lack of sufficient autologous tissues available for reconstruction necessitates the use of nongenital tissues for repair (Jordan, 1999). This challenge is further complicated by the need for multiple surgeries to achieve the goal of an aesthetically acceptable and functionally normal penis. Accomplishing this goal is crucial in maintaining self-esteem and preventing psychological scars in males (Spyropoulos et al., 2005; Lumen et al., 2008). Non-genital autologous tissues have been used since the 1930s, when phallic reconstruction was initially attempted using rib cartilage as a stiffener in patients with traumatic penile loss (Bogoraz, 1936; Goodwin et al., 470

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1952; Frumpkin, 1994). This method, which involved multiple, staged surgeries, was soon discouraged due to unsatisfactory functional and cosmetic results (Beheri, 1969; Frumpkin, 1994). Silicone rigid prostheses were popularized in the 1970s and have been used widely (Small et al., 1975; Bretan, 1989). However, biocompatibility issues have become problematic in selected patients that received the silicone implants (Kardar and Pettersson, 1995; Nukui et al., 1997). Recently, tissue transfer techniques using flaps from various non-genital sources – such as the groin, dorsalis pedis and osteocutaneous free fibula flap – have also been used for genital reconstruction (Jordan, 1999; Sengezer et al., 2004). However, operative complications – such as infection, graft failure and donor site morbidity – remain as limiting factors for this technique. In recent years, researchers in the field of tissue engineering have endeavored to develop biological substitutes that could be used to restore and maintain the normal function in many organ systems (Atala, 2003, 2006; Cross et al., 2003). This was based on the concept of cell transplantation as an approach to build viable tissues. Cell-based therapies using tissue engineering and regenerative medicine strategies have presented alternative possibilities for repairing pathological penile conditions (Atala, 2004). These experiments have led to novel methods of cell culture that allow autologous tissue to be grown ex vivo from a small sample of a patient’s own tissue. These techniques could be extremely useful in genital reconstruction, as they may be able to generate the large amounts of tissue required for a reconstruction. In addition, using engineered autologous tissue may be preferable to earlier methods because the engineered tissue would be biocompatible. In this chapter, we will focus on the development of tissue engineering approaches that are designed to repair genital tissue and restore normal function.

23.2

Basic principles of penile tissue engineering

The basic components required to engineer functional tissues are cells, scaffolds and in vivo environment. Tissue engineering strategies involve the use of biocompatible matrices either with or without cells. The matrices are either used as supporting scaffolds to promote and enhance tissue regeneration, or as cell delivery vehicles. When cells are used, donor tissue is dissociated into individual cells, which are expanded in culture, attached to a support matrix and introduced into the body to form functional tissues. This general concept has been demonstrated repeatedly in many tissue engineering applications. However, special considerations have to be given for each tissue or organ type due to the heterogeneity of function. Similarly, tissue engineering of penile tissue requires understanding of cellular, anatomical and functional characteristics.

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The penis houses three separate cylindrical structures, consisting of a paired corpora cavernosa and the corpus spongiosum which sits in the groove created by the corpora cavernosa (Belman, 1978). The corpus spongiosum encases the urethra, which serves as a conduit for the passage of urine and semen. The corpus cavernosum, also known as the erectile tissue, is composed of interconnected sinusoids surrounded by smooth muscle. Each of these cylinders is encased in a tough, thick sheath known as the tunica albuginea. Corpus cavernosum possesses unique tissue architecture which controls the inflow of blood to maintain an erectile state. In the flaccid state, the smooth muscle in the cavernosum is tonically contracted. This allows only a small amount of arterial blood flow to enter the sinusoidal space of corpus cavernosa. During sexual stimulation, nitric oxide (NO) released from nerve endings and the endothelium causes smooth muscle relaxation which permits influx of blood into the cavernosal sinusoids (Ignarro et al., 1990). This causes the penis to lengthen and widen rapidly until the capacity of the tunica albuginea is reached. The expanded sinusoidal walls press against one another and against the tunica albuginea encasing the corpora, which leads to occlusion of venous outflow, thus maintaining the rigid state of the penis during erection (Lue, 2000). These unique anatomical and functional features of the corporal tissue are specialized and require full understanding in order to engineer tissue that would function like native penile tissue. An essential component that is necessary to build penile tissue is cells. The main cell types that are required to achieve erectile function are smooth muscle and endothelial cells. These cells interact with each other in a coordinated manner to control erectile state and both of these cell types are necessary to build corporal tissue that could function normally. Thus, isolation and culture expansion of these cells are the initial steps toward building erectile tissue. These cells can be found in various tissues in the body. However, it is uncertain whether cells obtained from non-erectile tissues would display similar characteristics to the cells found in the corporal tissue. Studies have used various tissue sources to obtain smooth muscle and endothelial cells, including cavernosal tissue, and some studies obtained endothelial cells from human foreskin tissue (Kwon et al., 2002; Hu et al., 2006; Giuliani et al., 2008). Recently, stem and progenitor cells have been used to derive these cell types (De Coppi et al., 2007a, 2007b). Regardless of the cell sources used to obtain these cells, it is important to demonstrate that cellular characteristics are functionally normal. Another important component that needs to be considered is scaffold design. The scaffolds for corporal tissue engineering should be able to provide structural integrity, bear similar architecture, exhibit similar material characteristics, safely accommodate cells within the matrix and be biocompatible and biodegradable. A variety of synthetic and naturally derived

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materials have been studied to determine the ideal support materials for the regeneration of urological tissue (Tachibana et al., 1985; Atala et al., 1995; Oberpenning et al., 1999; El-Kassaby et al., 2003). Biodegradable synthetic materials, such as poly-lactic and glycolic acid polymers, have been shown to provide good support for growing cells in many models, and they can be easily fabricated and configured in a controlled manner. On the other hand, naturally derived materials, such as collagen and fibronectin, are biocompatible and provide a similar extracellular matrix environment to normal tissue (Olsen et al., 1992; Kropp et al., 1998; Atala et al., 1999). Recently, corporal tissue matrix that possesses similar tissue architecture to native corpora has been developed by removing all cellular components from donor corporal tissue (Kwon et al., 2002; Falke et al., 2003). This scaffolding system has been demonstrated to be useful in the engineering of functional corporal tissues. One of the technical difficulties in engineering corporal tissue may be the efficient delivery of cells to target locations within the scaffold. Cell seeding on to a scaffold and arranging them into a three-dimensional orientation that is identical to as delicate a structure as the native penis may be challenging. In order to maximize the delivery of cells into corporal scaffolds, a dynamic seeding method has been used effectively (Eberli et al., 2008). While manual cell seeding on to a scaffold may work for some tissue applications, in other tissues that require precise delivery of multiple cell types this process cannot be performed manually. Recent advances in technology development have allowed the use of inkjet printing techniques to deliver different cell types precisely to their designated target locations within a scaffold (Boland et al., 2006). Utilization of these technological modalities may allow for building improved tissues for functional restoration.

23.3

Engineering of functional corporal tissue

One of the major tissue components of the phallus is the corpus cavernosa. These are sponge-like tissues that hold blood during a normal erection. The ability to create autologous functional and structural corporal tissue de novo would be beneficial for treating various pathological conditions. To achieve this goal, tissue engineering efforts have been directed toward creating corpus cavernosal tissue for reconstructive procedures. Initial experiments have been performed in our laboratory to determine the feasibility of creating corporal tissue using cultured human corporal smooth muscle cells seeded on to biodegradable polymers (Kershen et al., 1998). Primary normal human corpus cavernosal smooth muscle cells were cultured, seeded on to the biodegradable polymer scaffolds and implanted subcutaneously in athymic mice. Implants were retrieved at 7, 14 and 24 days after surgery for analyses. Corporal smooth muscle cell layers were

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observed along the surface of the polymers at all time points. At the periphery of the implants, early vascular ingrowth was evident by 7 days. By 24 days post-implantation, there was evidence of polymer degradation. Smooth muscle phenotype was confirmed immunocytochemically and by Western blot analyses with antibodies specific to alpha-smooth muscle actin. This study provided the first evidence that cultured human corporal smooth muscle cells, in conjunction with biodegradable polymers, have the potential to create corpus cavernosum tissue in the laboratory. One major hurdle in the engineering of tissues and organs is the need to supply the new tissue with oxygen and nutrients. This can be achieved through adequate vascularization to the implanted tissue. Adequate vascular supply is especially important in the penis, as it is required for normal erectile function. Our laboratory investigated the possibility of developing vascularized human corporal tissue in vivo by combining smooth muscle and endothelial cells (Park et al., 1999). Primary human cavernosal smooth muscle cells and human endothelial cells (ECV 304) were seeded on biodegradable polymers. A total of 80 polymer scaffolds (60 seeded with cells and 20 without cells) were implanted in the subcutaneous space of athymic mice. Mice were killed at various time points up to 42 days after implantation. At retrieval, all polymer scaffolds that had been seeded with cells had formed distinct tissue structures and maintained their pre-implantation size, while the control scaffolds without cells had decreased in size. Histologically, all of the retrieved polymers seeded with corporal smooth muscle and endothelial cells contained surviving cells. Importantly, the presence of penetrating native vasculature was observed 5 days after implantation. In addition, the formation of multilayered strips of smooth muscle adjacent to endothelium was evident by 7 days after implantation. Increased smooth muscle organization and accumulation of endothelium lining the luminal structures were evident 14 days after implantation. A well-organized construct, consisting of muscle and endothelial cells, was noted at 28 and 42 days after implantation. There was no evidence of tissue formation in the controls. Immunocytochemical analyses using anti-von Willebrand factor (vWF) and anti-pan cytokeratins distinguished the origin of the vascular structures in each of the constructs. Anti-alpha-actin antibodies confirmed the smooth muscle phenotype. Smooth muscle fibers were progressively organized with time. The muscle to endothelial tissue ratio was approximately equivalent to the ratio of muscle and endothelial cell seeding prior to implantation (2 : 1). These experiments were the first to demonstrate that angiogenesis and vascularization in an engineered tissue can be facilitated by adding endothelial cells to the construct. They showed that corporal smooth muscle cells and endothelial cells seeded on biodegradable polymer scaffolds are able

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to form vascularized cavernosal muscle when implanted in vivo. These findings indicate that endothelial cells are able to act in concert with the native vasculature to accelerate angiogenesis in the implant. This study suggested that the creation of well-vascularized, autologous corporal-like tissue may soon be possible. Although tissue composed of corporal cells was successfully engineered in vivo, corporal tissue structurally identical to the native corpus cavernosum could not be achieved, due to the type of polymers used. We developed a naturally derived acellular corporal tissue matrix that possesses the same architecture as native corpora. In a study, we explored the feasibility of developing corporal tissue (consisting of human cavernosal smooth muscle and endothelial cells) in vivo, using the acellular corporal tissue matrices as a cell delivery vehicle (Falke et al., 2003). Acellular corporal tissue matrices processed from donor rabbit corpora were seeded with primary human cavernosal smooth muscle and endothelial cells. Tissue matrices seeded with cells and control matrices without cells were implanted in the subcutaneous space of athymic mice and followed for up to 8 weeks after implantation. Additional acellular tissue matrices seeded with cells were grown in culture for up to 4 weeks. Scanning electron microscopic examination of the seeded cells in vitro demonstrated a uniform attachment on the sinusoidal walls within the corporal tissue matrix. The implanted corporal tissue matrices maintained the seeded cells on the sinusoidal wall and showed host cell infiltration 3 days after implantation. Formation and migration of neovasculature into the sinusoidal spaces was evident by 1 week after implantation. Increasing organization of smooth muscle and endothelial cells lining the sinusoidal wall was observed at 2 weeks and continued with time. The corporal tissue matrices were completely covered with the appropriate cell architecture 4 weeks after implantation. This study demonstrated that human cavernosal smooth muscle and endothelial cells seeded on acellular corporal tissue matrices are able to form vascularized corporal structures in vivo. The use of these tissue matrices as cell delivery scaffolds allowed for the development of adequate constructs. The formation of corporal structures, similar to those of the native erectile tissue, may provide an additional armamentarium in the management of complex penile reconstructive challenges. In order to explore the possibility of replacing the entire cross-sectional segment of both corporal bodies with autologous engineered tissues, a study in rabbits was performed (Kwon et al., 2002; Chen et al., 2005). Acellular corporal collagen matrices were obtained by decellularizing donor rabbit penile tissue, and autologous corpus cavernosal muscle and endothelial cells were harvested, expanded and seeded on these matrices. The seeded constructs were then surgically implanted into the rabbit phallus. Functional and structural parameters – including cavernosography, cavernosometry,

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mating behavior, and sperm ejaculation – were observed for 6 months following implantation. The engineered corporal bodies demonstrated intact structural integrity on cavernosography and decreased maximal intracavernosal pressure on cavernosometry compared with non-surgical controls. Mating activity in the animals with engineered corpora appeared normal by 1 month after implantation. The animals with implanted corpora were placed with female rabbits, and copulation attempts were made within 30 seconds. Control animals that received acellular matrix alone showed a decrease in mating activity. In addition, the rabbits with implants were able to ejaculate, and the presence of sperm was confirmed in all rabbits with engineered corpora. The female rabbits that mated with the males implanted with engineered corpora conceived and delivered healthy pups. On gross examination, the corporal implants with cells showed continuous integration of the graft into native tissue. On histological examination, sinusoidal spaces and vessel walls lined with endothelium and smooth muscle were observed in the engineered grafts. Grafts without cells contained fibrotic tissue and calcifications with sparse corporal elements. These series of studies demonstrate that penile corpora tissue can be engineered and that this tissue can be sufficient for erection, mating, ejaculation and conception. However, further studies must be performed before these techniques can be used clinically, including observation of the longterm functionality of the neo-corpora, and the development of cell/matrix combinations with architecture that is nearly identical to that of native human corpus cavernosum. Although smooth muscle and endothelial cells are the major components of erectile tissue, other structures, such as connective tissue and nerves, are needed to achieve corpora that are structurally and functionally normal.

23.4

Engineered penile prosthesis

Although silicone is an accepted biomaterial for penile prostheses, biocompatibility remains a concern for selected patients. The use of a natural prosthesis composed of autologous cells may eliminate the concerns associated with silicone prosthesis. Initial studies have shown that chondrocytes suspended in biodegradable polymers can form cartilage structures when implanted in vivo (Atala et al., 1993, 1994). Thus, a feasibility study for creating natural penile prostheses made of cartilage was performed (Yoo et al., 1998). Chondrocytes harvested from the articular surface of calf shoulders were isolated and expanded in culture. The cells were seeded on to preformed cylindrical polyglycolic acid polymer rods (1 cm in diameter and 3 cm in length). The cell–polymer scaffolds were implanted in the subcutaneous space of athymic mice. The rods were retrieved at 1, 2, 4 and 6 months post-implantation. At retrieval, all polymer scaffolds seeded with cells had

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formed milky-white, rod-shaped solid cartilaginous structures and maintained their pre-implantation size and shape. The control scaffolds without cells failed to form cartilage. In addition, there was no evidence of erosion, inflammation or infection in any of the implanted cartilage rods. Compression, tension and bending studies showed that the cartilage structures were readily elastic and could withstand high degrees of pressure. Biomechanical analyses showed that the engineered cartilage rods possessed the mechanical properties required to maintain penile rigidity. Histological examination showed the presence of mature and well-formed cartilage in all the chondrocyte–polymer implants. The polymer fibers were progressively replaced by cartilage in time. In a subsequent study, the feasibility of applying the engineered cartilage rods in situ was investigated (Yoo et al., 1999). Autologous chondrocytes harvested from rabbit ear were grown and expanded in culture. The cells were seeded on to biodegradable poly-l-lactic acid-coated polyglycolic acid polymer rods and implanted into the corporal spaces of rabbits. The animals were killed at 1, 2, 3 and 6 months after implantation. Gross examination at retrieval showed the presence of well-formed, milky-white cartilage structures within the corpora at 1 month and there was no evidence of erosion or infection in any of the implant sites. Histological analyses with alcian blue and toluidine blue staining demonstrated the presence of mature and well-formed chondrocytes in the retrieved implants. This technology appears to be useful for the creation of autologous penile prostheses. The possibility of engineering human cartilage rods for use as penile prostheses has also been investigated (Kim et al., 2002). Chondrocytes were isolated from human ear via small biopsy. The cells were seeded on to rodshaped biodegradable polymer scaffolds (1.2 cm in diameter, 6.0 cm long), and these were maintained in bioreactors for 1 month. Subsequently they were implanted into the subcutaneous space of athymic rats. The mechanical properties of the engineered prostheses were compared with those of silicone prostheses. Human chondrocytes seeded on to polymer scaffolds formed milky-white cartilaginous rods of the same size as the initial implants. The engineered human cartilaginous rods were flexible, elastic and able to withstand high degrees of compressive forces. The mechanical properties were comparable with those of commercially available silicone prostheses. This study demonstrates the feasibility of creating human cartilage rods with a large dimension. In the future, the use of bioengineered prostheses could be applied to patients undergoing penile surgery for either congenital or acquired conditions. This autologous cartilage tissue could be used without corporal tissue for penile reconstruction or intracorporally for erectile dysfunction. Further studies of long-term tissue interaction, survival and maintenance are being conducted.

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23.5

Reconstruction of the tunica albuginea

Penile conditions – such as Peyronie’s disease, tumor resection or augmentation surgery – often require reconstruction of the tunica albuginea. A variety of materials have been used as grafts to achieve this purpose. These materials include skin, vein patch grafts, pericardium, tunica vaginalis, Dacron and Goretex (Leungwattanakij et al., 2001; Gholami et al., 2003; Dean and Lue, 2004). Schultheiss et al. generated a tissue engineered graft from fibroblasts seeded on to acellular porcine small intestine submucosa. However, it remains unclear whether this engineered graft will counteract or even enhance fibrosis in the graft area (Schultheiss, 2004; Schultheiss et al., 2004). Acellular bladder matrix is a collagen-based biomaterial produced from bladder lamina propria. It is biocompatible and possesses good characteristics for tissue handling. This matrix has previously been used for urethral and bladder reconstruction (Chen et al., 1999; De Filippo et al., 2002). Our laboratory has examined the applicability and functional outcome of the use of acellular bladder matrix for tunica repair. Biocompatibility testing was performed on the matrix, including assays for cell viability, mitochondrial metabolic activity and apoptosis. Next, approximately 50% of the dorsal tunica albuginea was excised in rabbits and the defect was replaced with an acellular matrix patch. Cavernosometry and cavernosography were performed. Cavernosometry of the repaired animals demonstrated normal intracavernosal pressures with visual erections. Cavernosography of the repaired corpora showed a normal anatomical configuration. The animals were killed 1, 2 and 3 months after surgery for analyses. Biomechanical analysis of the retrieved matrices demonstrated similar tensile strengths to native tunica. Histologically, there was only a minimal inflammatory response, which gradually decreased over time. These results show that acellular bladder matrix is biocompatible, durable and effective when used as a tunica substitute. The matrix may be useful as an off-the-shelf biomaterial for patients requiring tunica albuginea repair, but further studies are necessary before the technology can be applied in humans (Eberli et al., 2007).

23.6

Summary and future trends

The evolution of the field of tissue engineering has led to the development of possible new approaches for the reconstruction of genital tissues. In this chapter, various tissue engineering and regenerative medicine approaches for the creation of functional penile tissue in vivo were presented. In the future, small penile tissue biopsies could be obtained through minimally invasive techniques under local anesthesia. Autologous cells derived from

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the biopsy could be expanded and seeded on an appropriate scaffold material, leading to the development of a biocompatible replacement organ. However, despite major advances in this field, reconstruction of the whole penile structure and integration of the engineered organ into the circulatory and nervous systems, and initiation of host neuronal control are still challenges that must be met before true penile replacement is a reality.

23.7

Acknowledgement

The authors thank Dr Jennifer Olson for editorial assistance.

23.8

References

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of amniotic stem cell lines with potential for therapy’, Nat Biotechnol, 25, 100–106. de coppi p, callegari a, chiavegato a, gasparotto l, piccoli m, taiani j, pozzobon m, boldrin l, okabe m, cozzi e, atala a, gamba p and sartore s (2007b), ‘Amniotic fluid and bone marrow derived mesenchymal stem cells can be converted to smooth muscle cells in the cryo-injured rat bladder and prevent compensatory hypertrophy of surviving smooth muscle cells’, J Urol, 177, 369–376. dean r c and lue t f (2004), ‘Peyronie’s disease: advancement in recent surgical techniques’, Curr Opin Urol, 14, 339–343. eberli d, susaeta r, yoo j j and atala a (2007), ‘Tunica repair with acellular bladder matrix maintains corporal tissue function’, Int J Impot Res, 19(6), 602–609. eberli d, susaeta r, yoo j j and atala a (2008), ‘A method to improve cellular content for corporal tissue engineering’, Tissue Eng Part A, 14, 1581–1589. el-kassaby a w, retik a b, yoo j j and atala a (2003), ‘Urethral stricture repair with off-the-shelf collagen matrix’, J Urol, 169, 170–173. de filippo r e, yoo j j and atala a (2002), ‘Urethral replacement using cell seeded tubularized collagen matrices’, J Urol, 168, 1789–1792. falke g, yoo j j, kwon t g, moreland r and atala a (2003), ‘Formation of corporal tissue architecture in vivo using human cavernosal muscle and endothelial cells seeded on collagen matrices’, Tissue Eng, 9, 871–879. frumpkin a p (1994), ‘Reconstruction of male genitalia’, Am Rev Sov Med, 2, 14. gholami s s, gonzalez-cadavid n f, lin c s, rajfer j and lue t f (2003), ‘Peyronie’s disease: A review’, J Urol, 169, 1234–1241. giuliani s (2008), ‘Tissue engineering of the reproductive system’, in Atala A, Lanza R, Nerem R and Thomson J A (Eds) Principles of Regenerative Medicine, Burlington, Elsevier, pp. 1144–1148. goodwin w e and scott w w (1952), ‘Phalloplasty’, J Urol, 68, 903. hsu g l (2004), ‘Anatomy of the human penis: The relationship of the architecture between skeletal and smooth muscles’, J Androl, 25, 426. hu x, jiang z and liu n (2006), ‘A novel approach for harvesting lymphatic endothelial cells from human foreskin dermis’, Lymphat Res Biol, 4, 191–198. ignarro l j, bush p a, buga g m, wood k s, fukuto j m and rajfer j (1990), ‘Nitric oxide and cyclic GMP formation upon electrical field stimulation cause relaxation of corpus cavernosum smooth muscle’, Biochem Biophys Res Commun, 170, 843–850. jordan g h (1999), ‘Penile reconstruction, phallic construction, and urethral reconstruction’, Urol Clin North Am, 26, 1–13. kardar a and pettersson b a (1995), ‘Penile gangrene: a complication of penile prosthesis’, Scan J Urol Nephrol, 29, 355. kershen r t, yoo j j, moreland r b, krane r j and atala a (1998), ‘Novel system for the formation of human corpus cavernosum smooth muscle tissue in vivo’, J Urol, 159(suppl.), 156–160. kim b s, yoo j j and atala a (2002), ‘Engineering of human cartilage rods: potential application for penile prostheses’, J Urol, 168, 1794–1797. kropp b p, ludlow j k, spicer d, rippy m k, badylak s f, adams m c, keating m a, rink r c, birhle r and thor k b (1998), ‘Rabbit urethral regeneration using small intestinal submucosa onlay grafts’, Urology, 52, 138–142.

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kwon t g, yoo j j and atala a (2002), ‘Autologous penile corpora cavernosa replacement using tissue engineering techniques’, J Urol, 168, 1754–1758. leungwattanakij s, bivalacqua t j, reddy s and hellstrom w j (2001), ‘Long-term follow-up on use of pericardial graft in the surgical management of Peyronie’s disease’, Int J Impot Res, 13, 183–187. lue t f (2000), ‘Erectile dysfunction’, N Engl J Med, 342, 1802. lumen n, monstrey s, selvaggi g, ceulemans p, de cuypere g, van laecke e and hoebeke p (2008), ‘Phalloplasty: a valuable treatment for males with penile insufficiency’, Urology, 71, 272–276. nukui f, okamoto s, nagata m, kurokawa j and fukui j (1997), ‘Complications and reimplantation of penile implants’, Int J Urol, 4, 52. oberpenning f, meng j, yoo j j and atala a (1999), ‘De novo reconstruction of a functional mammalian urinary bladder by tissue engineering, Nat Biotechnol, 17, 149–155. olsen l, bowald s, busch c, carlsten j and eriksson i (1992), ‘Urethral reconstruction with a new synthetic absorbable device. An experimental study’, Scand J Urol Nephrol, 26, 323–326. park h j, yoo j j, kershen r t, moreland r and atala a (1999), ‘Reconstitution of human corporal smooth muscle and endothelial cells in vivo’, J Urol, 162, 1106–1109. ralph d j, garaffa g and garcia m a (2006), ‘Reconstructive surgery of the penis’, Curr Opin Urol, 16, 396–400. schultheiss d (2004), ‘Regenerative medicine in andrology: tissue engineering and gene therapy as potential treatment options for penile deformations and erectile dysfunction’, Eur Urol, 46, 162–169. schultheiss d, lorenz r r, gabouev a i, schlote n, wefer j and mertsching h (2004), ‘Functional tissue engineering of autologous tunica albuginea: a possible graft for Peyronie’s disease surgery’, Eur Urol, 45, 781–786. sengezer m, ozturk s, deveci m and odabasi (2004), ‘Long-term follow-up of total penile reconstruction with sensate osteocutaneous free fibula flap in 18 biological male patients’, Plast Reconstr Surg, 114(2), 439–450. small m p, carrion h m and gordon j a (1975), ‘Small-Carrion penile prosthesis: new implant for management of impotence’, Urology, 5, 479. spyropoulos e, christoforidis c, borousas d, mavrikos s, bourounis m and athanasiadis s (2005), ‘Augmentation phalloplasty surgery for penile dysmorphophobia in young adults: considerations regarding patient selection, outcome evaluation and techniques applied’, Eur Urol, 48, 121–127. tachibana m, nagamatsu g r and addonizio j c (1985), ‘Ureteral replacement using collagen sponge tube grafts’, J Urol, 133, 866–869. yoo j j, lee i and atala a (1998), ‘Cartilage rods as a potential material for penile reconstruction’, J Urol, 160, 1164–1168. yoo j j, park h j, lee i and atala a (1999), ‘Autologous engineered cartilage rods for penile reconstruction’, J Urol, 162, 1119–1121.

24 Tissue engineering in reproductive medicine A. S O P H O N S R I T S U K and C. E. B I S H O P, Wake Forest Institute For Regenerative Medicine, USA

Abstract: Reproduction is a fundamental feature of all known life. In humans, the reproductive period of women, which begins at puberty and lasts until menopause, is a complex part of reproductive biology. However, a variety of congenital and pathological conditions involving the female reproductive organs can interfere with normal function, and the field of reproductive medicine has grown in response to these conditions. Increasingly, physicians and scientists look to regenerative medicine and tissue engineering techniques to develop novel methods of genital reconstruction for the treatment of such disorders. Using these techniques, functional tissues such as uterus and ovary have been created, and such discoveries provide hope to women who previously could not conceive as a result of defects in the reproductive tract. In this chapter, we present protocols used for tissue engineering of the vagina, uterus and ovarian tissues, and discuss the clinical relevance of these techniques. Key words: reproduction, reproductive medicine, uterus, vagina, ovary.

24.1

Tissue engineering of the vagina

The vagina is a thin-walled, distensible, fibromuscular tube that extends from the vestibule of the vulva to the uterus. Histologically, the vagina is composed of four distinct layers. The mucosa consists of a stratified, nonkeratinized squamous epithelium. The next layer is the lamina propria, or tunica; it is composed of fibrous connective tissue. The muscular layer contains an inner circular layer and an outer longitudinal layer. The fourth layer consists of cellular areolar connective tissue containing a large plexus of blood vessels (Katz 2007). A variety of pathological conditions of the vagina require medical intervention. For example, cloacal malformation with vaginal abnormalities, uterovaginal agenesis resulting from Mayer–Rokitansky–Kuster–Hauser syndrome (MRKH syndrome) and trauma to the female genitalia can require numerous reconstructive surgeries. In addition, transexualism (Baytekin et al. 2007; Chin et al. 2007; Kwun et al. 2003) also requires construction of female genital organs. Gynecological malignancies are being 482

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diagnosed earlier, and treatment has been improved through the use of adjuvant chemotherapy and radiation. This has created a larger surviving population, many of whom require reconstructive procedures to repair the tissues damaged by the malignancy (Pusic and Mehrara 2006). In addition to treating these primary pathological conditions through reconstructive procedures, preservation of sexual function must be considered. The use of vaginal dilators to create a vagina or augment an existing structure that is inadequate for sexual intercourse is the current treatment of choice. Although success rates of up to 90% have been published, failure of the treatment in some oncological patients and disadvantages such as inconvenience and unpleasant sensation have been reported (Davies and Creighton 2007). Another option is surgical reconstruction, which traditionally involves the creation of a neovaginal space that is lined with either a split-thickness skin graft or a section of intestine. Surgical reconstruction is a major procedure with significant risks and complications (Davies and Creighton 2007). Surgical reconstruction with engineered vaginal tissue, rather than skin grafts or intestinal flaps, might be a promising future treatment. At our institute, we successfully demonstrated that vaginal epithelial and smooth muscle cells are able to reconstitute vaginal tissue in vitro (De Filippo et al. 2003). Recently, the creation of a vagina using autologous in vitro cultured vaginal tissue in a patient with MRKH syndrome has been reported (Panici et al. 2007).

24.2

Methods of vaginal tissue reconstitution

24.2.1 Cell isolation and culture In our experience, vaginal tissue can be reconstituted from a harvested vaginal sample that is 1 cm2 in size (De Filippo et al. 2003; Panici et al. 2007). Separation of the muscle and epithelial tissues is performed by microdissection under loop magnification. Epithelial and smooth muscle tissues are expanded in separate cultures. Isolation of the individual cell types involves one of two processes that consist of either an explant method or enzymatic digestion (De Filippo and Atala 2002). Explant method The separated epithelial or muscle tissues are minced into small pieces and these are individually placed on to culture dishes, where they dry and adhere to the surface. The pieces of tissues are incubated with the appropriate medium at 37 ⬚C in 5% CO2 until the progenitor cells in the sample form tissue islets. This usually takes approximately 5–7 days. The explants can be

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24.1 Vaginal epithelial cells (keratinocytes) growing out of a piece of vaginal tissue using the explant method. Scale bar represents 100 μm.

removed by gentle suction and the cells maintained with scheduled replacement of the medium every 24–48 h (De Filippo and Atala 2002) (Fig. 24.1). Enzymatic digestion The fastidious nature of epithelial cells sometimes makes the growth of large quantities of cells difficult. Therefore we have used enzymatic digestion to isolate vaginal epithelial cells with good success. Powder forms of collagenase type IV and dispase, a neutral protease, are combined and suspended with approximately 25 ml of keratinocyte serum-free medium (K-SFM; Life Technologies). This collagenase-medium solution is then filtered to ensure sterility. The vaginal tissue is chopped into several large pieces, immersed into the enzymatic solution and vigorously shaken for 25–30 min at 37 ⬚C. With gentle pipette suction, the cell-fluid suspension is transferred to another 50 ml tube and centrifuged for 5 min. Finally, the supernatant is removed and the cell pellet is resuspended in medium and distributed into culture dishes (De Filippo and Atala 2002; De Filippo et al. 2003). Seeding the cells on to collagen type IV-coated culture plates might enhance the growth of the cells (Panici et al. 2007).

24.2.2 Culture medium A number of commercially manufactured culture media are available for both epithelial and smooth muscle cell growth. We prefer Dulbecco’s modi-

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fied Eagle’s medium supplemented with 10% fetal bovine serum (DMEM/ FBS) for smooth muscle cells. For culture of vaginal epithelial cells, serumfree medium specifically for keratinocytes, supplemented with bovine pituitary extract and epidermal growth factor (K-SFM), is used (De Filippo and Atala 2002; De Filippo et al. 2003).

24.2.3 Subculture of the cells In order to pass the cells, the medium are removed and the cells are washed with phosphate-buffered saline (PBS). The cells are then incubated with 0.05% trypsin-EDTA (ethylenediaminetetraacetic acid) until they separate from the culture dish. The trypsin solution is neutralized with K-SFM, containing 10% FBS. After centrifugation, the cell pellet is resuspended with fresh medium and cells are plated on to fresh dishes. Each cell type is expanded to achieve a desired cell density of 10 × 106 cells/cm3 for epithelial cells and 20 × 106 cells/cm3 for smooth muscle cells (De Filippo et al. 2003).

24.2.4 Cell characterization Both types of cells are characterized by traditional immunohistochemistry. Epithelial cells and smooth muscle cells are identified with monoclonal anti-pancytokeratins AE1/AE3 and α-actin antibodies, respectively.

24.2.5 Scaffold construction and cell seeding We choose polyglycolic acid (PGA) as the scaffold for tissue formation in vivo. It is coated with a 50 : 50 copolymer of poly (dl-lactide-co-glycolide). The scaffolds are sterilized with ethylene oxide gas and pre-wetted with medium 24 h prior to seeding with cells. Vaginal epithelial and smooth muscle cells are seeded on to opposite sides of the scaffold and incubated at 37 ⬚C in 5% CO2 for 24–48 h. Finally, the seeded scaffolds are implanted into the animals.

24.2.6 Histology, molecular, tensile strength and organ bath studies When cell-seeded scaffolds are implanted into animals, it is possible to compare histology and protein expression between native vaginal tissue and the engineered vaginal structures with hematoxylin and eosin staining (Fig. 24.2) and Western blot analysis. Likewise, physiological and mechanical properties of the reconstituted vaginal tissue can be evaluated with organ bath studies and tensile strength testing.

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E

M

24.2 A hematoxylin and eosin stained section of a PGA scaffold seeded with vaginal epithelial and smooth muscle cells. This construct was implanted in vivo and retrieved 6 weeks later (E, epithelial cells; M, muscle cells; arrow demarcates the transitional layer; reduced from ×100) (De Filippo et al. 2003).

24.3

Tissue engineering of the uterus

Over the past few decades, there have been drastic improvements in infertility treatment. It is estimated that, in the Western world, between 8 and 30% of couples of reproductive age are infertile (Marchbanks et al. 1989; Thonneau et al. 1991). In approximately 40% of these couples, the cause is attributed to the woman (Hull et al. 1985). In some of these cases, the cause is related to the uterus itself. Congenital anomalies (such as a congenital absence of the uterus) or acquired anomalies (such as uterine fibroids, adenomyosis and carcinoma) all contribute to infertility rates. Unfortunately, many of these uterine abnormalities are not completely treatable. Since most couples want their own genetic children, in vitro fertilization and gestational surrogacy is the treatment of choice for couples dealing with uterine infertility. However, there are a variety of controversial issues surrounding surrogacy. For example, the potential psychological risks to a child born by surrogacy are not known. In addition the degree of involvement that the host may wish to have with the child following birth may not be acceptable to the parents, and there is the possibility that the host may wish to retain the child after birth. The lack of clear regulations regarding these issues limits widespread use of surrogacy (Brinsden 2003). For these reasons, surrogacy is even prohibited in some countries (van den Akker 2007). Uterine transplantation has long been studied in many animal models, such as the dog (O’Leary et al. 1969; Scott et al. 1970), cynomolgus monkey

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(Scott et al. 1971), sheep (Zhordania and Gotsiridze 1964) and rabbit (Confino et al. 1986). Few pregnancies were achieved in these studies (Brannstrom et al. 2003). The two major hurdles in these cases were the lack of adequate vascular supply for the transplanted uterus and the rejection of the graft (Brannstrom et al. 2003). While the latter issue could be resolved by immunosuppressive drugs, the purpose of a uterine transplant is to restore fertility, and these drugs have adverse effects on implantation and fetal development and should be avoided during pregnancy. Tissue engineering may provide new hope for those with uterine infertility. We have investigated the feasibility of a near-total uterine replacement using a fabricated autologous neo-uterus in a rabbit model. The tissueengineered uterus showed no signs of rejection and quickly became vascularized. It developed highly organized cellular, tissue and anatomical structures and demonstrated physiological and biomechanical responses similar to those of a normal uterus. Importantly, the neo-uterus was able to support natural conception, implantation, placentation and gestation to term, and resulted in the first birth of a normal live rabbit pup.

24.4

Methods of uterine tissue reconstitution

In all mammals, the uterus develops as a specialization of the paramesonephric or mullerian duct, which gives rise to the infundibula, oviducts, uterus, cervix and anterior vagina (Mossman 1987). The mature uterine wall is comprised of two functional compartments, the myometrium and endometrium. The myometrium is the smooth muscle layer of the uterine wall which consists of an inner circular layer derived from the intermediate layer of ductal mesenchymal cells and an outer longitudinal layer derived from the subperimetria mesenchyme. The endometrium is the inner mucosal lining of the uterus. It is derived from the inner layer of ductal mesenchyme. Histologically, the endometrium consists of two layers, the functional layer and the basal layer. The functional layer is adjacent to the uterine cavity. This layer is completely shed during menstruation. The basal layer, adjacent to the myometrium and below the functional layer, is not shed at any time during the menstrual cycle, and from it the functional layer develops (Gray et al. 2001). It has been proposed that stromal stem/progenitor cells reside in human endometrium (Gargett and Chan 2006). However, the study of uterine stem cells is hampered by the lack of specific markers and tests that could be used to identify stem cells. There have been reports that circulating stem cells of extrauterine origin, such as bone marrow stem cells, can differentiate into human endometrial tissue and contribute to endometrial regeneration (Du and Taylor 2007; Taylor 2004). In addition, recent studies utilized the label-retaining cell (LRC) technique as a functional approach for identifying adult stem cells in vivo. The technique revealed epithelial

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LRCs, constituting 3% of mouse endometrial epithelial cells, in the luminal epithelium rather than basal glands. Stromal LRCs were identified in 6% of stromal cells. Some of these were found in close association with luminal epithelium, while 40% were near the endometrial–myometrial junction (Cervello et al. 2007; Chan and Gargett 2006; Szotek et al. 2007). Such cells may be useful for reproductive tissue engineering applications, and to test this we have created an animal model. In our experience, New Zealand white rabbits have served as excellent animal models. The rabbit uterus is comprised of two separate uterine horns with no uterine body. Each uterine horn opens directly into the vagina. In addition, rabbits are induced ovulators and do not have an estrus cycle. This allows cell and tissue harvest techniques to be performed at any time, and ensures that the tissues will be uniform. Expansion of myometrial cells in culture for tissue engineering purposes has been previously achieved (Atala et al. 1993). We have also developed a cell culture system that allows for the rapid expansion of endometrial cells in large quantities (Baez and Atala 2002). Endometrial and myometrial cells can be obtained from uterine biopsy or hysterectomy specimens. These specimens should be transferred immediately to transport medium (Dulbecco’s modified Eagle’s medium with Ham’s F12 nutrient medium (DMEM/F12)). Biopsies exceeding 2 cm in diameter will remain viable in this medium for up to 3 days at 4 ⬚C.

24.4.1 Protocols for isolation and culturing of endometrial cells Enzymatic digestion and mechanical separation 1

Remove the specimen from the transport medium with sterile forceps and place in a 6-well tissue culture dish. 2 Wash the specimen with three applications of 5 ml sterile PBS containing 1% penicillin and streptomycin. 3 Use curved scissors to cut and expose the endometrial cavity. 4 Incubate with 0.1% collagenase type IV at 37 ⬚C for 40 min. 5 Scrape the endometrial layer by scalpel blade. 6 Centrifuge the collagenase solution at 80 × g for 5 min. 7 Remove supernatant and resuspend in culture medium. 8 Plate the cells on to tissue culture dish. Explant method (alternative option) 1

Remove the specimen from the transport medium with sterile forceps and place in a 6-well tissue culture dish.

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Wash the specimen with three applications of 5 ml sterile PBS containing 1% penicillin and streptomycin. Orient the specimen with the epithelial surface down and use curved iris scissors to dissect the submucosa from the muscle. A thin opaque white epithelial strip should be left after this procedure. Mince the epithelial strip finely, into very small fragments. Place the small fragments in a 10-cm culture dish and drop medium on to the small fragments. Change the medium every 2–3 days.

24.4.2 Protocols for culturing smooth muscle cells Enzymatic digestion 1

Mince the remaining muscle tissue from the endometrial cell isolation procedure finely, into very small fragments. 2 Incubate with 0.1% collagenase type I at 37 ⬚C for 2 h. 3 Separate smooth muscle cells from undigested tissue by filtering through a 100 μm cell strainer. 4 Add 15 ml of medium and centrifuge at 80 × g for 5 min. 5 Remove supernatant and resuspend with medium. 6 Count the cells, using a counting chamber. 7 Plate cells at a density of 2 × 104 cells/cm2. Explant method (alternative option) 1 2 3

Mince the remaining muscle tissue from the endometrial cell isolation procedure into very small fragments. Place the small fragments in a 10-cm culture dish and drop medium on to them. Change the medium every 2–3 days.

24.4.3 Culture medium We achieve an ample number of epithelial cells when using DMEM/F-12 medium (1 : 1 v/v) supplemented with epidermal growth factor (EGF) (5 ng/ ml), bovine pituitary extract (40 ng/ml) and 10% FBS. We prefer DMEM supplemented with 10% FBS for smooth muscle cells (Fig. 24.3).

24.4.4 Cell characterization Both types of cells are characterized by traditional immunohistochemistry. Epithelial cells are identified for epithelial antigens with monoclonal

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24.3 (a) Primary passage of endometrial cells isolated from uterine tissue. (b) Smooth muscle cells isolated from uterine tissue. Scale bar repre 100 μm.

antibodies to pancytokeratins AE1/AE3 and CD10 (Galabova-Kovacs et al. 2004; Hashizume et al. 2003; Sugawara et al. 1997). Nuclear hormone receptors can be detected with antibodies to estrogen and progesterone (Galabova-Kovacs et al. 2004; Kyo et al. 2003; Stiemer et al. 1995). Smooth muscle cells are identified with α-actin antibodies (Accinni et al. 1983; Al-Matubsi et al. 2001; Blin et al. 1996; Galabova-Kovacs et al. 2004; Rush et al. 2001).

24.4.5 Cell seeding and construct formation In order to develop a uterine construct in vitro, the expanded autologous myometrial and endometrial cells are seeded in a stepwise fashion on to the surface and lumen, respectively, of hollow, uterine-shaped scaffolds formed from PGA. Biodegradable polymer meshes of PGA are configured into the appropriate shape using 5-0 absorbable sutures. They are then coated with poly-dl-lactide-co-glycolide in chloroform (5%w/v) in order to increase stiffness and maintain structural integrity. The scaffolds are sterilized with ethylene oxide gas and pre-wetted with medium 24 h prior to seeding with cells. First, smooth muscle cells are seeded at a concentration of 6 × 107 cells/ml on to the outer side of the prefabricated scaffold, and the construct is cultured for 3 days in DMEM supplemented with 10% FBS. The epithelial cells are then seeded at a concentration of 6 × 107 cells/ml on the inner side of the matrix and cultured for an additional 48 hours prior to implantation. After a total of 5 days in vitro, the neo-uteri are transplanted back into the animals from which the cells were derived (Fig. 24.4).

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24.4 Left panel depicts an unseeded PGA/PLGA construct. Right panel depicts a PGA/PLGA construct seeded with autologous epithelial cells on the luminal side and myometrial cells on the outer surface prior to implantation in vivo.

24.4.6 Histology, molecular, tensile strength and organ bath studies In order to compare native uterine tissue and the in vitro reconstituted uterine tissues, hematoxylin and eosin staining and Western blot analyses can be performed on samples derived from both tissues. Likewise, physiological and mechanical properties of the reconstituted uterine tissue can be evaluated with organ bath studies and tensile strength testing.

24.5

Tissue engineering of the ovarian tissue

24.5.1 Folliculogenesis In humans, primordial germ cells that will form ovarian tissue are derived from the endoderm, and these cells differentiate in the yolk sac approximately 3 weeks after fertilization. The germ cells are mitotically active in the yolk sac and remain mitotically active while migrating to the genital ridge. The genital ridge is formed from mesenchyme and coelomic epithelium at about 4 weeks of gestation and will ultimately provide somatic cells to the follicle. At about 7 weeks, the primordial germ cells reach the genital ridge, and at about 16 weeks of gestation, mesenchyme cells from the differentiating ovarian stroma of the genital ridge condense and envelop the oogonia, forming the primordial follicle. The formation of primordial follicles is complete by about 18 weeks of gestation (Smitz and Cortvrindt 2002). The formation of the primordial follicle, which is characterized by a single layer of flattened somatic cells around the oogonium, terminates mitosis and initiates meiosis within the oogonia. Meiosis proceeds in the primordial follicles to the diplotene stage of prophase I, at which stage the condensed chromosomes are stored within the germinal vesicle of the

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oocyte. Once all follicles initiate meiosis, no further mitosis of the oogonia, now termed primary oocytes, can occur. Hence, the number of female gametes is set definitively at that stage of life. At this stage of follicular development, the average human fetus has 6–8 million germ cells. The numbers then rapidly decline to 2–4 million by birth, and at menarche, only about 400 000 remain in the ovary (De Pol et al. 1997; Smitz and Cortvrindt 2002). Continued growth of the oocyte, expansion of the granulosa cells to two to seven layers and differentiation of the theca layers of the granulosa cells mark the development of the secondary follicle. The secondary follicle has a diameter of approximately 100 μm. Late secondary follicles, or preantral follicles, have a diameter of approximately 150 μm. Preantral follicles are also characterized by the formation of the theca externa and by the initiation of vascularization of the follicle (Smitz and Cortvrindt 2002). When an oocyte reaches a mature size of about 120 μm and the number of granulosa cells increases to 2000–3000, a fluid-filled antrum forms within the follicle. Growth of the preantral follicles is follicle stimulating hormone (FSH)independent, whereas growth of the early antral stage follicles is FSHdependent. The oocyte undergoes cytoplasmic and nuclear maturation, and gap junctions form between the granulosa cells and oocyte. The gap junctions facilitate bidirectional communication between the oocyte and the granulosa cells (Eppig 2001; Eppig et al. 1997). The oocyte plays a dominant role in follicular maturation. Likewise, signals from granulosa cells are necessary to support oocyte development (diZerega and Hodgen 1981; Hillier 1994).

24.5.2 Cryopreservation of ovarian tissue The effect of chemotherapy and radiotherapy on fertility is a major concern to patients of reproductive age who are diagnosed with cancer. There are many options offered to these patients. In vitro fertilization (IVF) and subsequent embryo cryopreservation is the most successful, as pregnancy rates following embryo cryopreservation are reasonable. Oocyte cryopreservation would be an ideal technique for fertility preservation; however, the pregnancy rates following this technique are low. Human ovarian tissue banking is proposed as a method of preserving female fertility and offers the potential of restoring normal ovarian function and natural fertility. The main advantage of this technique is that no ovarian stimulation is required and thus the procedure can be performed on an urgent basis. Moreover, small immature follicles in the ovarian cortex probably withstand cryopreservation better than mature ones (Hovatta et al. 1996; Newton et al. 1996; Oktay et al. 1997). Cryopreservation of ovarian tissue rich in primordial and primary follicles is a strategy for oocyte banking. The rationale is

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to cryopreserve immature follicles within the ovarian tissue, before oocytes resume nuclear maturation (Revel and Schenker 2004). There are two potential methods to resume fertility from cryopreserved ovarian tissues. Ovarian autotransplantation in humans was described almost a century ago (Revel and Schenker 2004). Recently, a successful pregnancy resulting from transplantation of cryopreserved ovarian tissue has been reported (Demeestere et al. 2007). However, in patients with ovarian cancers, transplantation should not be performed since this may result in the transfer of malignant cells back to the patients (Revel and Schenker 2004). In these cases, in vitro folliculogenesis is another option. It entails harvesting mature oocytes in vitro from freeze–thawed ovarian cortex by isolating small follicles (Smitz and Cortvrindt 1999) or oocytes (Abir et al. 1997) from the surrounding stroma and growing them to maturity. Oocyte in vitro maturation (IVM) is currently feasible only in the latest stages of follicular development and requires optimization before it can be used in the clinic. Freshly aspirated GV stage (prophase I) oocytes can be matured successfully in the laboratory to re-initiate and complete the first meiotic division to metaphase II. This has resulted in pregnancies and live births in polycystic ovary syndrome patients (Cha and Chian 1998). However, the limited number of GV stage oocytes in the ovary is a concern. A technique that could be used to mature earlier stage oocytes in vitro is urgently needed for the recovery of cryopreserved ovarian tissue.

24.5.3 Tissue engineering of the ovarian follicle Communication between the various ovarian cell types and the stroma is required for normal follicle development. The extracellular matrix (ECM) within the follicle is believed to play a role in regulating follicle development (Woodruff and Shea 2007). Three-dimensional culture systems for ovarian tissue have been developed, and these mimic in vivo conditions so that the importance of the ECM can be studied. Cell-to-cell communication can be disrupted in the flat architecture created in a two-dimensional system, but cells that aggregate on particular ECM surfaces or within threedimensional hydrogels maintain cytoplasmic processes and gap junctions, allowing for proper communication between neighboring cells (Amsterdam et al. 1989; Ben-Rafael et al. 1988; Ben-Ze’ev and Amsterdam 1986; Richardson et al. 2000). One three-dimensional tissue engineering approach used to explore this phenomenon is encapsulation of the follicles (Heise et al. 2005). Microencapsulation has been used to provide structural support for a variety of tissues such as pancreatic islets (de Groot et al. 2004) and thyroid follicles (Glaser et al. 1999). Microencapsulation using alginate has been investigated using the oocytes of several animal species, including human (Torre

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24.5 Ovarian follicle encapsulated in a calcium alginate bead. Scale bar represents 50 μm. Reproduced from Heise et al. (2005) with permission.

et al. 2006), and the results have been promising (Kreeger et al. 2006; Rowghani et al. 2004; Vigo et al. 2005; Xu et al. 2006) (Fig. 24.5). Alginate is one of the most commonly applied biomaterials for microencapsulation due to its biocompatibility, high affinity for water, and its ability to form gels under mild conditions in the presence of calcium ions (Amsden and Turner 1999; Smidsrod and Skjak-Braek 1990). Alginate is composed of chains of alternating blocks of mannuronic acid, which contribute to the elastic properties of the gel, and glucoronic acid, which provides mechanical strength, stability, porosity and gelling properties (Yang and Wright 1999; Zimmermann 1999). Alginates are extracted from all species of brown algae and contain differing compositions of mannuronic acid/mannuronic acid, mannuronic acid/glucoronic acid and glucoronic acid/glucoronic acid blocks, which creates variations in strength and stability. Alginate gel beads are reported to have a high porosity range and only limit the diffusion of large proteins (Heise et al. 2005). It has been reported that substances of molecular weight (MW) less than 2 × 104 – such as glucose, l-trytophan (MW = 204) and α-lac-tolalbumin (MW = 1.54 × 104) – are able to diffuse freely into and out of calcium alginate beads at approximately the same rate as in water, while larger proteins, such as albumin (MW = 6.9 × 104), could not diffuse freely into the calcium alginate beads. Although larger proteins (MW > 2 × 104) have difficulty diffusing into these beads, diffusion from the

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bead into a surrounding solution devoid of the substrate is not hindered until the molecular weight of the substance approaches 3 × 105(Heise et al. 2005; Smidsrod and Skjak-Braek 1990). Several studies using alginate microencapsulation have investigated the regulating factors that control follicle development, and these studies demonstrated interesting results. For example, the transition from a two-layered secondary follicle to a multilayered secondary follicle was promoted by collagen I and RGD (arginine – glycine – aspartic) peptide. The transition from multilayered secondary to antral follicle was FSH-dependent and appears to be delayed by laminin or fibronectin (Kreeger et al. 2006). Inclusion of FSH in the alginate bead restores the follicle growth response (Heise et al. 2005).

24.6

Method for culturing follicles

24.6.1 Follicle isolation The ovary is incubated in α-minimal essential medium (MEM) supplemented with 1% fetal calf serum (FCS), 0.1% type I collagenase and 0.02% DNaseI at 37 ⬚C, 5% CO2 for 30 min. 2 Follicles are mechanically isolated using insulin-gauge needles in L15 media containing 1% FCS under a dissecting microscope. 3 Follicles are maintained in α-MEM at 37 ⬚C, 5% CO2 for 2 h before encapsulation. 4 Follicles are evaluated and measured. Only follicles displaying the following characteristics during the incubation period are selected for encapsulation: (a) diameter corresponding to secondary or preantral follicles; (b) intact follicles with some attached, fibroblast-like thecal cells; and (c) follicles containing a visible, immature oocyte that is round and centrally located within the follicle. 1

24.6.2 Calcium alginate gel encapsulation Based on the methods of McGee et al. (Heise et al. 2005), gel encapsulation can be accomplished as follows. 1

Twenty to thirty follicles are transferred with a glass pipette to a solution of sodium alginate. 2 The mixture of follicles and sodium alginate is slowly released through a 25-gauge needle to form a thin stream. A stream of 0.2 μm-filtered air is positioned at the tip of the needle to cut the mixture stream into small droplets to create droplets with diameters between 250 and 400 μm. 3 The droplets are allowed to fall into a beaker containing a stirred solution of CaCl2 (0.1 M). The droplets immediately gel to form beads.

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Beads are removed from the beaker using glass pipettes and transferred to 12 × 75 mm polypropylene culture tubes containing media. Follicles are maintained at 37 ⬚C, 5% CO2 and the medium is changed every 48 h.

Based on the methods of Shea and Woodruff et al. (Kreeger et al. 2006), gel encapsulation can also be accomplished as follows. 1 2 3 4

Single follicles are pipetted into the middle of an alginate droplet (2–3 μl) that is suspended on a polypropylene mesh (0.1 mm opening). The mesh is immersed in sterile 50 mM calcium chloride for 2 min, and this causes each alginate droplet to gel into a bead. The mesh is rinsed in culture media and the beads are plated at one follicle per well in 96-well plates in 100 μl of culture media. Follicles are maintained at 37 ⬚C, 5% CO2. The fresh medium is replaced every 48 h.

24.6.3 Culture media α-MEM, supplemented with 10 mIU/ml bovine serum albumin (BSA), 1 mg/ml bovine fetuin, 5 μg/ml transferrin and 5 ng/ml selenium is used for follicle culture.

24.6.3 Oocyte recovery and evaluation 1 2 3 4

When needed, follicles are removed from the alginate beads by degrading the gel with 10 mIU/ml alginate lyase for 30 min at 37 ⬚C, 5% CO2. The released follicles are transferred to media. Granulosa cells are removed from oocytes. Oocytes are considered to be in metaphase I if neither the germinal vesicle nor the first polar body are visible. If a polar body is present in the perivitelline space, the oocytes are classified as metaphase II. Fragmented or shrunken oocytes are considered to be degenerated.

24.7

Conclusions

There have been numerous studies aimed at the creation of functional vaginal, uterine and ovarian tissues, but many problems still remain. Clinical trials involving large numbers of patients must be performed before tissueengineered vaginal constructs can be considered a mainstream method for female genital reconstruction. In the case of uterine tissue, additional studies of biomaterials and their biochemical properties must also be performed so that a method of producing adequately vascularized uterine tissue constructs can be developed. Additionally, the hormonal response of engi-

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neered endometrium must be studied, and these studies are ongoing in our laboratory. The process of ovarian cryopreservation and its use is dependent upon overcoming issues involved in the recovery of early stage cryopreserved oocytes. Finally, further understanding of the molecular biology, cell biology and developmental biology involved in the growth of female reproductive organs would help us to develop improved tissue engineering techniques that could restore both sexual and reproductive function in women.

24.8

Acknowledgements

The authors would like to Matthew Heise, Richard Koepsel, Alan J Russel and Elizabeth A McGee for permission to use figures. The authors would also like to thank Dr Jennifer Olson for editorial assistance with this manuscript.

24.9

References

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25 Regenerative medicine of the kidney N. G U I M A R A E S - S O U Z A, R. S O L E R and J. J. YO O, Wake Forest Institute for Regenerative Medicine, USA

Abstract: End stage renal failure is a devastating condition that involves multiple organ systems in affected individuals. Although management strategies, such as dialysis, are able to prolong survival, renal transplantation is the only definitive treatment that can restore complete kidney function. However, transplantation is limited by several factors, such as critical donor shortage, complications due to chronic immunosuppressive therapy and rejection. Recent advances in cell technologies have facilitated the development of cell-based approaches for kidney tissue regeneration. The kidney exhibits a complex cellular composition, which makes bioengineering renal tissue a challenging task. Isolation and expansion of specific types of renal cells, such as erythropoietin-producing cells, may be a good approach for selective cell therapy. In addition, different cell sources, improved growth environments, novel differentiation factors and the use of synthetic or natural biomaterials have led to exciting regenerative medicine strategies that may be used to restore renal function. Although these approaches hold promise, implementation of these technologies in the clinic is still distant. This chapter reviews current regenerative medicine and tissue engineering strategies that may be used to develop innovative alternatives to improve, restore or replace renal function. Key words: cell therapy, kidney, kidney failure, regenerative medicine, tissue engineering.

25.1

Introduction

The human kidney contains approximately 1.0–1.3 million nephrons that serve as the basic functional unit. These units are essential for maintaining body homeostasis. They ensure that electrolyte balance is maintained and that metabolic waste products are excreted. In addition, the kidney secretes hormones that regulate other tissue systems (Rose, 1987). In particular, it secretes hormones that affect the cardiovascular system and that are required for maintenance of proper hemodynamics (renin, angiotensin II, nitride oxide, endothelin and bradykinin). Importantly, the kidney also regulates red blood cell production in the bone marrow by secreting erythropoetin in response to changes in oxygen tension. It also regulates bone metabolism through maintenance of calcium and phosphorus concentra502

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tions in the body and through secretion of 1,25 dihydroxyvitamin D3. Other functions of the kidney include catabolism of peptide hormones and synthesis of glucose under fasting conditions (gluconeogenesis) (Anon., 2002). The functions of kidney are compromised when renal tissue is extensively damaged. Progression of kidney damage usually results in tubular atrophy and glomerulosclerosis, which eventually lead to loss of functional nephron units and fibrosis (Anon., 2002, Levey et al., 2005). Numerous patients worldwide suffer from kidney disease every year and a substantial number of these patients progress to end stage renal disease (ESRD) (Jones et al., 1998). ESRD is a devastating condition that involves multiple organ systems in affected individuals (Anon., 2002, Culleton et al., 1999, Drey et al., 2003, Eknoyan et al., 2004, Levey et al., 2005; McCullough, 2002, Tveit et al., 2002). Although management strategies, such as dialysis, can prolong survival via filtration of the blood, other kidney functions are not replaced, thus leading to long-term consequences such as anemia and malnutrition (Amiel et al., 2000a, Barreto et al., 2008, Chazan et al., 1991, Cohen et al., 1994, Moe et al., 2005) Currently, renal transplantation is the only definitive treatment that can restore complete kidney function – including filtration, production of erythropoietin and production of 1, 25 dihydroxyvitamin D3. However, transplantation is limited by several factors, such as critical donor shortage, complications due to chronic immunosuppressive therapy and graft failure (Amiel et al., 2000a, Chazan et al., 1991, Cohen et al., 1994, Ojo et al., 2000, Ojo et al., 2001, Port et al., 2004, Wolfe et al., 1999). The limitations of current therapies for renal failure have led investigators to explore the development of alternative therapeutic modalities that could improve, restore or replace renal function. The emergence of cell-based therapies using tissue engineering and regenerative medicine strategies has presented alternative possibilities for the management of pathological renal conditions (Amiel and Atala, 1999, Amiel et al., 2000a, Amiel et al., 2006, Humes and Szczypka, 2004, Kurian et al., 1999, Perin et al., 2007). The concept of kidney cell expansion and cell transplantation using tissue engineering and regenerative medicine techniques has been proposed as a method to augment either isolated or total renal function. Despite the fact that the kidney is one of the more challenging organs to regenerate due to its complexity, investigative advances made to date have been promising (Amiel and Atala, 1999, Kurian et al., 1999, Perin et al., 2007).

25.2

Basic components of renal tissue engineering

25.2.1 Cells The structure of the kidney includes numerous arterioles, capillaries, glomeruli and tubules that interconnect in a specific, three-dimensional pattern

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to filter the blood as it passes through the organ and to excrete waste through a collecting system. In addition, nephrons and collecting ducts are composed of multiple functionally and morphologically distinct segments. This complex tissue structure has been difficult to replicate as more than 30 different cell types are present in the kidney and work in concert to maintain normal function. Thus, regenerative medicine techniques attempt to provide appropriate conditions for the long-term survival, differentiation and growth of many cell types. Because cells are considered the basic building blocks, recent efforts have focused on the search for a reliable cell source for renal tissue regeneration. In addition, optimal growth conditions have been extensively investigated to provide adequate enrichment for achieving stable renal cell expansion systems (Carley et al., 1988, Horikoshi et al., 1988, Humes and Cieslinski, 1992, Milici et al., 1985, Schena, 1998). Isolation of particular cell types that produce specific factors, such as erythropoietin, may also be a good approach for selective cell therapies. However, total renal function would not be achieved using this approach. In order to create kidney tissue that would deliver full renal function, a culture containing all of the cell types that constitute the functional nephron units should be used. Optimal culture conditions to nurture renal cells have been extensively studied and the cells grown under these conditions have been reported to maintain their cellular characteristics (Lanza et al., 2002). Furthermore, when these cultured renal cells were placed in a threedimensional culture environment, they were able to reconstitute renal structures. The search for a reliable cell source has also been expanded to stem and progenitor cells. The use of these cells for tissue regeneration is attractive due to their ability to differentiate and mature into specific cell types required for regeneration. This is particularly useful in instances where primary renal cells are unavailable due to extensive tissue damage. Bone marrow-derived human mesenchymal stem cells have been shown to be a potential source and have the ability to differentiate into several cell lineages (Ikarashi et al., 2005; Kale et al., 2003, Prockop, 1997). These cells have been shown to participate in kidney development when they are placed in a rat embryonic niche that allows for continued exposure to a repertoire of nephrogenic signals (Yokoo et al., 2005). These cells, however, were found to contribute mainly to regeneration of damaged glomerular endothelial cells after injury. The major cell source for kidney regeneration was found to originate in the intrarenal cells in an ischemic renal injury model (Ikarashi et al., 2005, Lin et al., 2005). Recently, the isolation of a side population of cells in adult human kidney was described. These authors previously showed that side population cells from mouse kidney can differentiate into multiple lineages (Inowa et al., 2008). Circulating stem cells have also been shown to transform into tubular and glomerular epithelial cells, podocytes,

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mesangial cells and interstitial cells after renal injury (Gupta et al., 2002, Ito et al., 2001, Iwano et al., 2002, Kale et al., 2003, Lin et al., 2003, Poulsom et al., 2001, Rookmaaker et al., 2003). Another promising stem cell source for renal tissue regeneration is human amniotic fluid-derived stem cells (hAFSCs). These cells are capable of contributing to the development of primordial kidney structures including the renal vesicle and C- and S-shaped bodies (Perin et al., 2007). These observations suggest that controlling stem and progenitor cell differentiation could lead to successful regeneration of kidney tissues.

25.2.2 Biomaterials Although isolated renal cells are able to retain their phenotypic and functional characteristics in culture, transplantation of these cells in vivo may not result in structural remodeling that is appropriate for kidney tissue. In addition, cell or tissue components cannot be implanted in large volumes due to limited diffusion of oxygen and nutrients (Folkman and Hochberg, 1973). Thus, a cell-support matrix, preferably one that encourages angiogenesis, is necessary to allow diffusion across the entire implant. A variety of synthetic and naturally derived materials has been examined in order to determine the ideal support structures for regeneration (Atala et al., 1995, Atala et al., 2006, El-Kassaby et al., 2003, Oberpenning et al., 1999, Tachibana et al., 1985). Biodegradable synthetic materials, such as poly-lactic and polyglycolic acid polymers, have been used to provide structural support for cells. Synthetic materials can be easily fabricated and configured in a controlled manner, which make them attractive options for tissue engineering. However, naturally derived materials – such as collagen, laminin and fibronectin – are more biocompatible and provide cells with an extracellular matrix environment that is similar to normal tissue. For this reason, collagen-based scaffolds have been used increasingly in many applications (Freed et al., 1994, Hubbell et al., 1991, Mooney et al., 1996, Wald et al., 1993).

25.3

Approaches for the regeneration of renal tissue

25.3.1 Developmental approaches Implantation of a kidney precursor such as the metanephros has been proposed as a possible method for achieving functional restoration. A recent study devised a method based on the stages of metanephric kidney development to engineer in vitro kidney-like tissue containing functional tubular transporters and glomeruli with apparent early vascularization. The Wolffian duct was isolated from timed-mated Holtzan rats at embryonic day 13, and induced to bud in vitro. Then, each isolated bud was induced to undergo

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branching. The branched ureteric buds were recombined with metanephric mesenchyme that had been isolated from rat kidney rudiments. After 4–6 days of mutual induction, the recombined tissue resembled a late-stage embryonic kidney, which was implanted into a host animal. After 14 days, renal structures such as glomeruli and evidence of early vascularization were observed (Rosines et al., 2007a). In one study, human embryonic metanephroi transplanted into the kidneys of an immune-deficient mouse model developed into mature kidneys (Dekel et al., 2003). The transplanted metanephroi produced urine-like fluid, but failed to develop ureters. This study suggests that development of an in vitro system in which metanephroi can be grown may lead to transplant techniques that produce a small replacement kidney within the host. In another study, the metanephros was divided into mesenchymal tissue and ureteral buds, and each of the tissue segments was cultured in vitro (Steer et al., 2002). After 8 days in culture, each portion of the tissue had grown to the size of the original mesenchymal tissue from which it was derived. A similar method was used for ureteral buds, which also propagated. These results indicate that if the mesenchyme and ureteral buds were placed together and cultured in vitro, a metanephros-like structure would develop and that the metanephros could be propagated under optimal conditions. Transplantation of metanephroi into a non-immunosuppressed rat omentum has also been performed. The implanted metanephroi are able to undergo differentiation and growth that is not confined by a tight organ capsule (Drey et al., 2003, Rogers et al., 1998). When metanephroi with an intact ureteric bud were implanted, the metanephroi enlarged and became kidney-shaped within 3 weeks. In addition, they were able to develop a well-defined cortex and medulla. Mature nephrons and collecting system components taken from these kidney-like structures were shown to be indistinguishable from those of normal kidneys by light and electron microscopy (Hammerman, 2002). Moreover, these structures became vascularized via arteries that originated at the superior mesenteric artery of the host (Hammerman, 2002). The transplanted metanephroi survived for up to 32 weeks post-implantation (Rogers et al., 1999). These studies show that the developmental approach may be a viable option for regenerating renal tissue for functional restoration. Controlling kidney differentiation is one of the key factors for successful tissue regeneration. In one study, hyaluronic acid was shown to possess the ability to simultaneously modulate ureteric bud branching, promote mesenchymal-to-epithelial transformation and promote differentiation of both metanephric mesenchyme and the ureteric bud depending on its concentration and molecular weight. These findings suggest that hyaluronic acid might be useful for creating a three-dimensional scaffold for in vitro kidney engineering and for promoting tubule regeneration in injured or

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cryopreserved kidneys, based on the biocompability and crosslinking capability of this compound (Rosines et al., 2007b). Some groups have also described the formation of specific kidney structures in vitro. In particular, a new method for creating kidney tubules has been reported. Using micromachined molds, channels were formed in extracellular matrix gels and were subsequently filled with Madin–Darby canine kidney epithelial cells. Upon cell adherence, a second layer of extracellular matrix gel was introduced. After gelation of the second gel layer, the samples were placed into a 12-well plate filled with culture media to allow the cells to assemble into tubular tissue. After 3–5 days, the epithelial cells self-assembled into tubular structures of up to 1 cm. These structures developed a lumen lined by a monolayer of polarized epithelial cells at 7 days. Kidney epithelial cells are characterized by their polarization, which is represented by the establishment of specific basolateral and luminal surfaces. The authors conclude that this method is feasible for generating kidney epithelial tubules with natural shape and dimensions in a short time (Schumacher et al., 2008). Another interesting method of creating kidney structures in vitro may optimize in vivo growth and differentiation of renal stem cells. Embryonic renal tissue containing stem and progenitor cells was mounted within a perfusion culture container at the interface of an artificial interstitium made from polyester. The space between the inner wall and the growing tissue was filled with an artificial interstitium made of a polyester fleece to reduce the dead volume space. During the entire culture period the fleece was in contact with the mounted tissue. In this way, the culture medium flowed through the fleece as in natural capillaries and ensured an equally distributed liquid exchange. Thus, the dead space volume was minimized and a constant fluid environment was created within the culture container. Additionally, mechanical protection was provided to the growing tissue. Renal tubules developed in chemically defined Iscove’s modified Dulbecco’s medium without serum addition and without coating with extracellular matrix proteins (Hu et al., 2007). In another study, non-genetically modified stem cells derived from human amniotic fluid were used to demonstrate the development and differentiation into de novo kidney structures during organogenesis in vitro (Perin et al., 2007). hAFSCs were isolated from human male amniotic fluid obtained between 12 and 18 weeks’ gestation. Green fluorescent protein and Lac-Z-transfected hAFSCs were microinjected into murine embryonic kidneys (12.5–18 days gestation) and were maintained in a special co-culture system in vitro for 10 days. The hAFSCs were characterized during their integration and differentiation in concert with the growing organ. The hAFSCs were capable of contributing to the development of primordial kidney structures including renal vesicle and Cand S-shaped bodies. Reverse transcriptase polymerase chain reaction confirmed expression of early kidney markers for zona occludens-1, glial-derived

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neurotrophic factor and claudin. This study shows that human amniotic fluid-derived stem cells may represent a potentially limitless source of ethically neutral, unmodified pluripotential cells for kidney regeneration.

25.3.2 Tissue engineering approaches The ability to grow and expand renal cells is one of the essential requirements for engineering tissues. The feasibility of achieving renal cell growth, expansion and in vivo reconstitution using tissue engineering techniques has been investigated (Atala et al., 1995). Rabbit kidney cells were grown, expanded and seeded on to biodegradable polyglycolic acid scaffolds. The cell-seeded scaffolds (those with proximal tubular cells, glomeruli, distal tubular cells and a mixture of all three cell types) were implanted subcutaneously into athymic mice. Animals were killed at 1 week, 2 weeks, and 1 month after implantation and the retrieved implants were analyzed. An acute inflammatory phase and a chronic foreign body reaction were seen, accompanied by vascular ingrowth by 7 days after implantation. Histological examination demonstrated progressive formation and organization of nephron segments within the polymer fibers with time. Renal cell proliferation in the cell–polymer scaffolds was detected by in vivo labeling of replicating cells with the thymidine analog bromodeoxyuridine (BrdU). BrdU incorporation into renal cell DNA was confirmed using monoclonal antiBrdU antibodies. These results demonstrated that renal-specific cells can be successfully harvested and cultured, and can subsequently attach to artificial biodegradable polymers. The renal cell–polymer scaffolds can be implanted into host animals where the cells replicate and organize into nephron segments. The polymer, which serves as a cell delivery vehicle, degrades with time. Initial experiments showed that implanted cell–polymer scaffolds gave rise to renal tubular structures. However, it was unclear whether the tubular structures reconstituted de novo from dispersed renal elements, or if they merely represented fragments of donor tubules that survived the original dissociation and culture processes intact. This was investigated further by Fung and colleagues (Fung et al., 1996). Mouse renal cells were harvested and expanded in culture. Subsequently, single isolated cells were seeded on biodegradable polymers and implanted into immune competent syngeneic hosts. Renal epithelial cells were able to reconstitute tubular structures in vivo. Sequential analyses of the retrieved implants over time demonstrated that renal epithelial cells first organized into a cord-like structure with a solid center. Subsequent canalization into a hollow tube could be seen after 2 weeks. Histological examination with nephron segment-specific lactins showed successful reconstitution of proximal tubules, distal tubules, loops of Henle, collecting tubules and collecting ducts. These results showed that

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single suspended cells are capable of reconstituting into tubular structures, with homogeneous cell types within each tubule. Although these studies demonstrate that renal cells seeded on biodegradable polymer scaffolds are able to form some renal structures in vivo, complete renal function could not be achieved in these studies. A subsequent study was designed in order to create a functional artificial renal unit that could produce urine. Mouse renal cells were harvested, expanded in culture and seeded on to a tubular device constructed from polycarbonate. The tubular device was connected at one end to a silastic catheter which terminated into a reservoir. The device was implanted subcutaneously in athymic mice. Histological examination of the implanted device demonstrated extensive vascularization as well as formation of glomeruli and highly organized tubule-like structures. Immunocytochemistry for osteopontin, which is secreted by proximal and distal tubular cells and the cells of the thin ascending loop of Henle, stained the tubular sections. Immunohistochemical staining for alkaline phosphatase stained proximal tubulelike structures. Uniform staining for fibronectin in the extracellular matrix of newly formed tubes was observed. Importantly, the reservoir of this device was filled with a yellow fluid upon retrieval. This fluid contained 66 mg/dl uric acid (as compared with 2 mg/dl in plasma) suggesting that the tubules were capable of unidirectional secretion and concentration of uric acid. The creatinine assay performed on the collected fluid showed an 8.2fold increase in concentration, as compared with serum. These results demonstrated that single cells from multicellular structures can become organized into functional renal units that are able to excrete high levels of solutes in a urine-like fluid (Yoo et al., 1996). In order to determine whether renal tissue could be formed using an alternative cell source, nuclear transplantation (therapeutic cloning) was performed to generate histocompatible tissues, and the feasibility of engineering autologous renal tissues in vivo using these cloned cells was investigated (Lanza et al., 2002). Nuclear material from bovine dermal fibroblasts was transferred into unfertilized, enucleated donor bovine oocytes. Renal cells from the cloned embryos were harvested, expanded in vitro and seeded on to three-dimensional scaffolds. These devices were implanted into the back of the same steer from which the cells were cloned, and were retrieved 12 weeks later. This process produced functioning renal units. Urine-like fluid production and viability were demonstrated after transplantation back into the nuclear donor animal. Chemical analysis of the excreted fluid suggested unidirectional secretion and concentration of urea, nitrogen and creatinine. Microscopic analysis revealed formation of organized glomeruli and tubular structures. Immunohistochemical and reverse transcriptionpolymerase chain reaction (RT-PCR) analysis confirmed the expression of renal mRNA and proteins. Finally, delayed-type hypersensitivity testing and

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in vitro proliferative assays showed that there was no rejection response to the cloned cells. This study indicates that the cloned renal cells were able to organize into functional tissue structures that were genetically identical to the host. Thus, the generation of immune-compatible cells using therapeutic cloning techniques is technically feasible and could be useful for the engineering of renal tissues for autologous applications. However, a naturally derived tissue matrix with existing threedimensional kidney architecture would be preferable to the synthetic materials that have been used in these studies. In addition to providing good biocompatibility, a naturally derived matrix would allow for transplantation of a larger number of cells, resulting in greater renal tissue volumes. Thus, we developed an acellular collagen-based kidney matrix, which is identical to the native renal architecture. We investigated whether these collagenbased matrices could accommodate large volumes of renal cells and form kidney structures in vivo. Acellular collagen matrices, derived from porcine kidneys, were obtained through a multi-step decellularization process. During this process, serial evaluation of the matrix for cellular remnants was performed using histochemistry, scanning electron microscopy (SEM) and RT-PCR. Mouse renal cells were then harvested, grown and seeded on 80 decellularized collagen matrices using a concentration of 30 × 106 cells/ ml. Forty cell–matrix constructs grown in vitro were analyzed 3 days, and 1, 2, 4 and 6 weeks after seeding. The remaining 40 cell-containing matrices were implanted in the subcutaneous space of 20 athymic mice. The animals were killed 3 days and 1, 2, 4, 8 and 24 weeks after implantation for analysis. Gross morphology, SEM, histochemical, immunocytochemical and biochemical analyses were performed. SEM and histological examination confirmed the acellularity of the processed matrix. RT-PCR performed on the kidney matrices demonstrated the absence of any RNA residues. Renal cells seeded on the matrix adhered to the inner surface and proliferated to confluency 7 days after seeding, as demonstrated by SEM. Histochemical and immunocytochemical analyses performed using hematoxylin and eosin, periodic acid Schiff, alkaline phosphatase, anti-osteopontin and anti-CD-31 identified stromal, endothelial and tubular epithelial cell phenotypes within the matrix. Renal tubular and glomerulus-like structures were observed 8 weeks after implantation. Methyl thiazolyl-tetrazolium (MTT) and tritiated thymidine incorporation assays performed 6 weeks after cell seeding demonstrated a population increase of 116% and 92%, respectively, as compared with the 2 week time points. This study demonstrates that renal cells are able to adhere to and proliferate on collagen-based kidney matrices. The renal cells reconstitute renal tubular and glomeruli-like structures in the kidney-shaped matrix. The collagen-based kidney matrix system seeded with renal cells may be useful in the future for augmenting renal function.

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Cell-based therapy for kidney disease

Augmentation of renal function without the use of an artificial device system would be preferable in some ways, as implantation procedures are invasive and may result in unnecessary complications. Our group has also investigated the feasibility of creating renal cell systems for in situ implantation within the native kidney tissue. Primary renal cells from 4-week-old mice were grown and expanded in culture. These renal cells were labeled with fluorescent markers, combined with a collagen gel and injected into mouse kidneys to attempt the in vivo formation of renal tissues. Collagen injection without cells and sham-operated animals served as controls. In vitro reconstituted renal structures and in vivo implanted cells were retrieved and analyzed. The implanted renal cells formed tubular and glomerular structures within the kidney tissue, as confirmed by the fluorescent markers. There was no evidence of renal tissue formation in the control and the sham-operated groups. These results demonstrate that single renal cells combined with a collagen gel-based scaffold are able to reconstitute kidney structures in vivo. The implanted renal cells appear to self-assemble into tubular and glomerular structures within the existing kidney tissue. Such a system may be the preferred approach to engineer functional kidney tissues for the treatment of ESRD. Although the concept of renal cell therapy has been demonstrated in several studies in which implanted, culture-expanded cells show the formation of renal structures, the efficiency of the process of structural reconstitution could not be assessed upon implantation in vivo. Reconstitution of renal structure during the culture expansion stage followed by implantation was proposed to provide a more controlled assessment of renal tissue in vivo (Joraku et al., 2005). A three-dimensional collagen-based culture system was developed to facilitate the formation of three-dimensional renal structures in vitro. After 1 week of growth, individual renal cells began to form renal structures resembling tubules and glomeruli. Histologically, these structures show phenotypic resemblance to native kidney structures. The reconstituted tubules stained positively for Tamm–Horsfall protein, which is expressed in the thick ascending limb of the loop of Henle and distal convoluted tubules. This study shows that renal structures can be reconstituted in a three-dimensional culture system, and this process may be used for renal cell therapy applications. While many studies have been conducted to achieve improvements in total function, investigations have also been directed toward augmenting selective kidney functions. In one study, cultured renal cells were examined to determine whether erythropoietin (EPO)-producing cells (Badylak et al., 2005) derived from kidney tissue could be used for the treatment of renal failure-induced anemia (Aboushwareb et al., 2008). Renal cells from 7- to

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10-day-old mice were culture expanded. The cells were characterized for EPO expression at each subculture stage. The levels of EPO expression were analyzed from renal cells incubated under normoxic and hypoxic conditions. The cultured renal cells expressed EPO at each subculture stage tested. This study indicates that EPO-producing cells may be used as a potential treatment option for anemia caused by chronic renal failure. There has also been considerable progress in the development of stem cell-based therapies for renal failure. For example, the existence of nontubular cells in adult mouse kidney that express stem cell antigen-1 (Sca-1) has been reported. This population of small cells includes a CD45-negative fraction that lacks hematopoietic stem cell lineage markers and resides in the renal interstitial space. In addition, these cells are enriched for β1-integrin, are cytokeratin negative and show minimal expression of surface markers that typically are found on bone marrow-derived mesenchymal stem cells. Clonally derived lines can be differentiated into myogenic, osteogenic, adipogenic and neural lineages. These renal Sca-1 cells were injected directly into the renal parenchyma of C57BL/6 wild-type mice, shortly after ischemic/reperfusion injury. After 1 month the injected cells had adopted a tubular phenotype and populated the renal tubule after ischemic injury. These adult kidney-derived cells may potentially contribute to kidney repair and may be important in the development of future regenerative medicine strategies (Dekel et al., 2003). Another study demonstrated that parietal epithelial cells (PECs) in the Bowman’s capsule exhibit coexpression of the stem cell markers CD24 and CD133 as well as expression of the stem cell-specific transcription factors Oct-4 and BmI-1. Lineage-specific markers are absent in this population. This population, which was purified from cultured encapsulated glomeruli, revealed self-renewal potential and a high cloning efficiency. Under appropriate culture conditions, individual clones of CD24+CD133+ PECs could be induced to generate mature, functional tubular cells with phenotypic features of proximal and distal tubules, osteogenic cells, adipocytes and neuronal cells. These cells were injected into severe combined immune-deficient (SCID) mice following rhabdomyolysis-induced acute renal failure. This treatment resulted in the regeneration of the tubular structures of different portions of the nephron, and it significantly ameliorated morphological and functional kidney damage. This study demonstrated the existence of resident multipotent progenitor cells in adult kidney and suggests a possible therapeutic use for these cells (Sagrinati et al., 2006).

25.5

Summary

The search for alternative treatment modalities for ESRD has been accelerated by the increasing demand for renal transplantation combined with the

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critical shortage of donor organs. Recent advances in cell technologies have allowed for development of cell-based approaches for kidney tissue regeneration. Various tissue engineering and regenerative medicine approaches that aim to achieve functional kidney support were presented in this chapter. Although these approaches are stimulating and hold promise, implementation of these technologies in the clinic is still distant. While it has been demonstrated that renal cells are able to reconstitute into functional kidney tissues in vivo and that the concept of cell transplantation is possible, numerous challenges must be met in order to translate these techniques into clinical therapies. Some of these challenges include: the generation of a large tissue mass that would augment systemic renal function; the integration of engineered renal tissue that includes adequate vascularization and excretory systems into the host; and the development of reliable renal failure model systems for testing the efficacy of cell-based technologies.

25.6

Acknowledgement

The authors thank Dr Jennifer Olson for editorial assistance.

25.7

References

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mooney, d. j., mazzoni, c. l., breuer, c., mcnamara, k., hern, d., vacanti, j. p. and langer, r. (1996) Stabilized polyglycolic acid fibre-based tubes for tissue engineering. Biomaterials, 17, 115–24. oberpenning, f., meng, j., yoo, j. j. and atala, a. (1999) De novo reconstitution of a functional mammalian urinary bladder by tissue engineering. Nat Biotechnol, 17, 149–55. ojo, a. o., hanson, j. a., meier-kriesche, h., okechukwu, c. n., wolfe, r. a., leichtman, a. b., agodoa, l. y., kaplan, b. and port, f. k. (2001) Survival in recipients of marginal cadaveric donor kidneys compared with other recipients and wait-listed transplant candidates. J Am Soc Nephrol, 12, 589–97. ojo, a. o., hanson, j. a., wolfe, r. a., leichtman, a. b., agodoa, l. y. and port, f. k. (2000) Long-term survival in renal transplant recipients with graft function. Kidney Int, 57, 307–13. perin, l., giuliani, s., jin, d., sedrakyan, s., carraro, g., habibian, r., warburton, d., atala, a. and de filippo, r. e. (2007) Renal differentiation of amniotic fluid stem cells. Cell Prolif, 40, 936–48. port, f. k., dykstra, d. m., merion, r. m. and wolfe, r. a. (2004) Organ donation and transplantation trends in the USA, 2003. Am J Transplant, 4 Suppl 9, 7–12. poulsom, r., forbes, s. j., hodivala-dilke, k., ryan, e., wyles, s., navaratnarasah, s., jeffery, r., hunt, t., alison, m., cook, t., pusey, c. and wright, n. a. (2001) Bone marrow contributes to renal parenchymal turnover and regeneration. J Pathol, 195, 229–35. prockop, d. j. (1997) Marrow stromal cells as stem cells for nonhematopoietic tissues. Science, 276, 71–4. rogers, s. a., lowell, j. a., hammerman, n. a. and hammerman, m. r. (1998) Transplantation of developing metanephroi into adult rats. Kidney Int, 54, 27–37. rogers, s. a., powell-braxton, l. and hammerman, m. r. (1999) Insulin-like growth factor I regulates renal development in rodents. Dev Genet, 24, 293–8. rookmaaker, m. b., smits, a. m., tolboom, h., van ‘t wout, k., martens, a. c., goldschmeding, r., joles, j. a., van zonneveld, a. j., grone, h. j., rabelink, t. j. and verhaar, m. c. (2003) Bone-marrow-derived cells contribute to glomerular endothelial repair in experimental glomerulonephritis. Am J Pathol, 163, 553–62. rose, b. d. (1987) Pathophysiology of Renal Disease, McGraw-Hill, New York. rosines, e., sampogna, r. v., johkura, k., vaughn, d. a., choi, y., sakurai, h., shah, m. m. and nigam, s. k. (2007a) Staged in vitro reconstitution and implantation of engineered rat kidney tissue. Proc Natl Acad Sci U S A, 104, 20938–43. rosines, e., schmidt, h. j. and nigam, s. k. (2007b) The effect of hyaluronic acid size and concentration on branching morphogenesis and tubule differentiation in developing kidney culture systems: potential applications to engineering of renal tissues. Biomaterials, 28, 4806–17. sagrinati, c., netti, g. s., mazzinghi, b., lazzeri, e., liotta, f., frosali, f., ronconi, e., meini, c., gacci, m., squecco, r., carini, m., gesualdo, l., francini, f., maggi, e., annunziato, f., lasagni, l., serio, m., romagnani, s. and romagnani, p. (2006) Isolation and characterization of multipotent progenitor cells from the Bowman’s capsule of adult human kidneys. J Am Soc Nephrol, 17, 2443–56. schena, f. p. (1998) Role of growth factors in acute renal failure. Kidney Int Suppl, 66, S11–5.

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schumacher, k. m., phua, s. c., schumacher, a. and ying, j. y. (2008) Controlled formation of biological tubule systems in extracellular matrix gels in vitro. Kidney Int, 73, 1187–92. steer, d. l., bush, k. t., meyer, t. n., schwesinger, c. and nigam, s. k. (2002) A strategy for in vitro propagation of rat nephrons. Kidney Int, 62, 1958–65. tachibana, m., nagamatsu, g. r. and addonizio, j. c. (1985) Ureteral replacement using collagen sponge tube grafts. J Urol, 133, 866–9. tveit, d. p., hypolite, i. o., hshieh, p., cruess, d., agodoa, l. y., welch, p. g. and abbott, k. c. (2002) Chronic dialysis patients have high risk for pulmonary embolism. Am J Kidney Dis, 39, 1011–17. wald, h. l., sarakinos, g., lyman, m. d., mikos, a. g., vacanti, j. p. and langer, r. (1993) Cell seeding in porous transplantation devices. Biomaterials, 14, 270–8. wolfe, r. a., ashby, v. b., milford, e. l., ojo, a. o., ettenger, r. e., agodoa, l. y., held, p. j. and port, f. k. (1999) Comparison of mortality in all patients on dialysis, patients on dialysis awaiting transplantation, and recipients of a first cadaveric transplant. N Engl J Med, 341, 1725–30. yokoo, t., ohashi, t., shen, j. s., sakurai, k., miyazaki, y., utsunomiya, y., takahashi, m., terada, y., eto, y., kawamura, t., osumi, n. and hosoya, t. (2005) Human mesenchymal stem cells in rodent whole-embryo culture are reprogrammed to contribute to kidney tissues. Proc Natl Acad Sci USA, 102, 3296–300. yoo, j. j., ashkar, s. and atala, a. (1996) Creation of functional kidney kidney structures with excretion of kidney-like fluid in vivo. Pediatrics, 98S, 605.

26 Stem cells and kidney regeneration S. S E D R A K YA N, L. P E R I N and R. E. D E F I L I P P O, Childrens Hospital Los Angeles, Keck School of Medicine University of Southern California, USA

Abstract: The role of stem cells in regenerative medicine is evolving rapidly. Scientists have looked at various sources of stem cells searching for alternative approaches in regenerative medicine. Acute renal failure (ARF) is a potentially devastating disorder in clinical medicine, with high mortality rates. Among many therapeutic options researchers have recently focused on stem cells with optimism that these pluripotential single units of life may hold the potential to overcome the current challenges in regenerative medicine. Several studies have addressed the roles of endogenous bone-marrow-derived stem cells in the repair of injured kidney. Exogenous stem cells with different origins have shown to contribute to structural and functional recovery of both glomerular and tubular compartments. Key words: kidney, acute renal failure, stem cells, regenerative medicine.

26.1

Introduction

The potential role of stem cells in regenerative medicine is under intensive investigation. A growing shortage of donor organs has heightened interest in developing novel strategies for tissue regeneration, such as stem cells, as an alternative to heterologous organ replacement. Over the past decade, scientists have looked at various sources of stem cells ranging from embryonic stem cells to adult stem cells, searching for alternative therapies in clinical medicine. Because stem cells have the capacity to differentiate into a variety of cell types, they are particularly appealing to tissue engineers because they represent a prospective inexhaustible reservoir of cells that can be used to regenerate new tissue for transplantation. Acute and chronic renal failure is a syndrome of permanently declining renal function subsequent to a variety of acute and gradual insults. Surgery and trauma are examples of acute injury, resulting into acute renal failure (ARF). When patients are younger or when the injury is less severe, renal tubules can regenerate and eventually return the kidney back to its normal function within days. However, in more severe cases or in older patients the damage to the kidney becomes an irreversible process with subsequent loss 518

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of function resulting in long-term dialysis and a marked increase in patient mortality. In contrast to ARF, chronic renal disease results from more unremitting causes, most commonly diabetes, but also hypertension, congenital malformations, autoimmune disorders, or chronic infection that can affect the individual for many years before organ failure is achieved. End stage renal disease (ESRD) usually occurs when kidney function is less than 10% of normal.1 The number of patients with ESRD is increasing annually throughout the world, and has reached epidemic proportions in the United States with an estimated 390 000 US residents currently receiving treatment. Currently, dialysis or allogenic transplantation are the only viable treatments despite significant complications with both. This, along with an increasing shortage of organs, has heightened interest in developing novel methods of therapy for kidney replacement. Pharmacological strategies have focused on targeting mechanisms and individual mediators thought to contribute to ischemic or toxic renal injury in experimental acute kidney injury (AKI) models. However, translational research efforts have yielded disappointing results.2–5 In the search for alternative approaches to dialysis and kidney transplantation scientists have recently focused on stem cells. Intensive research within the past decade has revealed the potential of stem cells as a therapeutic tool by virtue of the unique stem cell tropism and proregenerative capacity.6–8 It is noteworthy that stem cells undergo selfrenewal and differentiation not only in singeneic and allogenic situations, but also across the species barrier.9 The kidney is a complex organ with very important functions that are vital for the organism. These vital functions are made possible by specialized cell types that compose the glomeruli and tubules of the nephron within the surrounding extracellular matrix. Trying to identify an appropriate source of stem cells that will ultimately function to replace these specialized cells is very difficult. Both endogenous and exogenous sources of stem cells have been extensively studied by exploring their unique potential for the bioengineering of kidney cells or tissue. However, it is unclear now whether stem cells have the ability to recapitulate entirely the very complex differentiation pathways involved in kidney regeneration and completely replace one or all of the very complex cell types involved in this process. Therefore, this remains a very active area of research among many investigators today. In this chapter we will discuss stem cells and their possible role in kidney regeneration by focusing on the recent achievements in this area.

26.2

Endogenous stem cells

The regeneration capacity of a tissue is determined in part by whether it contains endogenous stem cells. Tissue-based stem cells have traditionally

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been viewed as multipotential precursor cells that are capable of generating tissue-specific differentiated cells.10 Organs such as the kidney, lung, liver, and heart possess these characteristics.11–14 Adult stem cells can only proliferate for a limited number of generations and their response to differentiation signals declines after each generation. In recent years, it has become clear that adult stem cells have remarkable plasticity to the extent that they can differentiate into lineages other than the tissue of origin.

26.2.1 Bone marrow-derived stem cells The potential role of mesenchymal stem cells (MSCs) as a tool for cellbased therapies aimed at kidney regeneration is an emerging interest among various scientific groups. The possibility that bone marrow (BM)-derived stem cells might contribute functionally to renal tubule regeneration is still a matter of debate. Several studies15–17 have demonstrated the presence of Y-chromosome-bearing cells in female kidneys transplanted into male recipients in humans as well as in in vivo models in mice, suggesting that stem cells from the male BM migrated and integrated into the transplanted female kidney. However, the number of integrated cells remains very low. In addition, Poulsom et al.15 demonstrated that BM-derived cells could contribute to the regeneration of the tubular epithelium after damage. Meanwhile, other studies have challenged the regenerative role of BMderived stem cells in kidney tubular epithelium, by putting more emphasis on renal progenitor cells.

26.2.2 Tissue-specific progenitor/stem cells Cells with the characteristics of renal progenitor cells have been isolated from the tubular fraction of normal renal cortex by targeting CD133, a marker known to be expressed by hematopoietic stem and progenitor cells, undifferentiated human intestine-derived epithelial cells, and embryonic kidney.18 Once isolated, these cells lack the expression of hematopoietic markers (CD34 and CD45), whereas they express some MSC markers (such as CD29, CD90, CD44, and CD73). Moreover, they express Pax2, an embryonic renal marker, suggesting their renal origin. The study by Lin et al.19 undermines the contribution of BM-derived stem cells in tubular regeneration by showing that within kidney tubules there exists a subset of cells that have the capability to proliferate rapidly after injury. In addition, a stem cell population in the papilla of the kidney also exists, regulated under a slow cell cycle during organ homeostasis that is induced toward rapid proliferation during injury.20 Further data provided by Humphreys et al.21 confirms the role of intrinsic epithelial cells in kidney repair after injury, by using genetic phase mapping techniques. According to their evidence,

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surviving tubular cells are the predominant cellular components involved in the repopulation of ischemic injured tubuli. Thus, terminally differentiated residual tubular cells may also contribute to kidney repair. Tubular epithelial cells that survive the damage are able to proliferate, dedifferentiate, and re-enter the cell cycle. During the dedifferentiation process they express early mesenchymal markers such as vimentin, and pax-2, a factor involved in kidney development. Moreover, it was confirmed that when enhanced green fluorescent protein (eGFP)-positive mature renal tubular epithelial cells where reactivated, they were able to proliferate at high rates and participate in tubular regeneration (almost 90%) after an ischemic injury to the kidney.22

26.2.3 Cell fusion – possible stem cell mechanism or not Methodological hindrances in tracking BM-derived cells in the injured kidney could be the possible reason for such contradictions in experimental results among different scientific groups. Cell fusion may also account for discrepancy in the different studies. Indeed, two studies23,24 have reported between 10 and 50% BM-derived stem cell engraftment in the kidney during regeneration after injury by fusion with resident tubular epithelial cells, rather than by transdifferentiation. Fang et al.23 reported up to 20% engraftment of BM-derived Stem cells in the kidney by fusion with resident epithelial cells during regeneration after folic acid injury. Moreover, significant replacement of damaged proximal tubular epithelium by fusion with tubular cells reaching up to 50%, was reported in fumarylacetoacetatehydrolase-deficient mice receiving BM transplants.24 This novel mechanism of transdifferentiation, which can occur in vivo, produces function cells in liver25 and brain.26 This is, however, a controversial issue, because in other experimental systems – skeletal muscle and pancreatic islets of langerhans – the cell fusion process has been excluded as a way of explaining BM stem cell plasticity.27,28 The origin of endogenous BM-derived stem cells that home in on the injured kidney remains to be elucidated. It has not been determined whether these cells homing in on the kidney are of hematopoietic or mesenchymal origin.

26.3

Exogenous stem cells

26.3.1 Embryonic stem cells Embryonic stem cells are derived from blastocysts and have generated vast interest in the scientific community owing to their pluripotential capacities. These cells propagate readily and remain undifferentiated in culture in the presence of leukemia inhibitory factor (LIF).29,30 When LIF is withdrawn,

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embryonic stem cells form aggregates called embryoid bodies (EB) that generate a variety of specialized cell types including neural, cardiac, and pancreatic cells.31 However, extraction of these cells has historically involved the destruction of embryos, with the associated highly controvertible ethical dilemmas. Kidney markers involved in the beginning of nephrogenesis are expressed during the early steps of EB formation, while terminally differentiated renal cell types are present in late EB development. Cells expressing markers characteristic of differentiated podocytes and epithelial cells of distal renal tubules were demonstrated within the EB.32 The same group also demonstrated that these cells are capable of resembling complex glomerular-like structures. Thus, embryonic stem cells show great potential, but safety concerns owing to their uncontrolled growth and tumorgenic properties, with a high propensity to develop proliferative abnormalities, has limited their ultimate practicality in clinical applications. In spite of this, there are reports of successful isolation of renal progenitors from embryonic stem cell-derived EBs by targeting T/GFP+, a marker known to be expressed during mesodermal specification – including intermediate mesoderm from which the kidney is derived – these renal progenitors integrate long-term into renal proximal tubules in vivo without teratoma formation.33

26.3.2 Bone marrow-derived stem cells Hematopoietic stem cells A number of studies have demonstrated that the administration of in vitro expanded stem cells may protect and reverse ARF. BM-derived hematopoietic stem cells (HSCs) and BM-derived MSCs have been extensively studied in the search for more effective treatment options that would give survival benefit, the key clinical outcome in ARF. Several studies demonstrated the participation of HSCs in renal tubular repair. Lin et al.16 showed that purified preparations of HSCs isolated from Rosa 26 mice, participated in kidney tubular repair when transplanted into female non-transgenic mice with ischemic acute renal injury. Another demonstration of such translineage differentiation by HSCs into renal epithelial cells was derived from the studies of Kale et al.,34 with evidence showing that HSCs, homing in on BM of irradiated mice, were mobilized by renal Ischemia–reperfusion (IR) into peripheral blood; the HSCs engrafted the kidney where they integrated and differentiated into tubular epithelium in the areas of damage. However, the HSC infusion and mobilization protocol has not been proven to be effective, failing to exert any protective effects; it has rather been associated with increased severity of renal tubular injury and mortality. In contrast, in a cisplatin-induced ARF model, treatment of mice with granulocyte-colony

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stimulating factor (G-CSF) mobilized Lin− CD34+ ckit+, and Sca+ cells significantly ameliorated renal function and reduced tubular necrosis35 by accelerating tubular regeneration and preventing apoptosis.36 Mesenchymal stem cells Based on experiments by Morigi et al.37 demonstrating that the beneficial effect of BM-derived stem cells is not ascribed to the HSCs but rather to the MSCs, all subsequent studies were focused on the therapeutic potential of MSCs. Administration of in vitro expanded MSCs has been shown to have beneficial effects in cisplatin-, glycerol- and IR-induced acute renal injury models.37–41 Further experiments provided evidence regarding the translineage differentiation potential and acquisition of epithelial phenotype of MSCs. Yokoo et al.42 injected MSCs from BM into kidneys during development and confirmed their integration into various compartments of the kidney, suggesting engraftment of these cells within nephron structures. It has been documented that MSCs, given 1 day after cisplatin, strongly protect from renal function impairment at days 4 and 5, as evaluated by blood urea nitrogen (BUN) assessment.43 In addition to engraftment and proliferation of MSCs to the site of tubular epithelial injury, it is suggested that recruitment of MSCs may also act by accelerating, to a remarkable extent, tubular cell proliferation. This has been demonstrated by experiments, in cisplatin- and glycerol-induced injury models, showing a high number of positive cells for Ki-67, a nuclear marker of cell proliferation, within the tubuli at the time at which renal function was ameliorated in stem cell-transplanted recipients.44 However, little is known about the molecular mechanisms that underlie MSC recruitment. Currently, several factors are recognized to have an important role in the migration of transplanted MSCs to sites of injury – such as chemokine receptor CXCR4, CD44, and possibly insulin-like growth factor-1 (IGF-1).38,43,44 Various studies have initiated the hypothesis that the functional benefits of MSCs could be a result of their ability to produce growth and trophic factors.45–48 Bonventre and co-workers,49 have emphasized in their studies that the process of renal epithelial cell replication occurs too rapidly for MSCs to truly transdifferentiate into tubular cells. In addition, the percentage of exogenous MSCs found in the tubules was less than 0.1% of the total population of injected cells at 24–48 h post-injection. Therefore, it is almost impossible that in such a short period of time these cells could play a predominant role in the repair of renal structures. Another important hypothesis that is under investigation is the possibility that MSCs may mediate their renoprotective effect by affecting the inflammatory process, implied by the renal upregulation of anti-inflammatory interleukin (IL)-10 – as well as organ protective growth factors fibroblast growth factor (βFGF),

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transforming growth factor (TGF-α), and antiapoptotic Bcl-2 – following acute renal injury. Reduced expression of proinflammatory cytokines such as IL-1β, tumor necrosis factor (TNF-α), and interferon-γ (IFN-γ), as well as inducible nitric oxide synthase (iNOS), is also in line with this new hypothesis of a possible paracrine involvement of MSCs in the pathogenesis of ARF.50

26.3.3 Amniotic fluid/placental stem cells More recently, Atala and his group51 have derived a novel stem cell population described as originating from human amniotic fluid that exhibits both embryonic and mesenchymal stem cell characteristics. These amniotic fluid stem cells (AFSCs) have extensive plasticity and are capable of differentiating into representative cell types of all three germ layers. The surface marker profile of AFSCs and their expression of the transcription factor Oct4 suggests that they represent an intermediate stage between pluripotent embryonic stem cells and lineage-restricted adult stem cells. They are easily expanded in culture while maintaining their pluripotential capacity. Moreover, unlike embryonic stem cells they do not form teratomas in vivo.51 Perin et al.52 have demonstrated that AFSCs, without any prior modifications, are able to integrate and proliferate when injected into embryonic kidneys in an in vitro co-culture system. The injected AFSCs were detected within primordial kidney, including in tubules and developing nephrons, confirmed by expression of early markers of kidney development (such as glial cell-derived neutrophic factor (GDNF), zona occluden-1 (ZO-1), and claudin). Some in vivo preliminary experiments were also performed in which direct injection of AFSCs into damaged kidneys showed that the cells were able to survive and integrate into tubular structures and expressed mature kidney markers after 3 weeks. Physiological parameter levels in these animals, which increased significantly after injury, were restored shortly after injection of AFSCs as previously demonstrated for BM-derived MSCs.53 It is therefore suggested from these results that AFSCs participate in similar immunological mechanisms to those postulated for MSCs, as discussed previously, during early phases of injury and perhaps contribute towards the eventual structural repair of the damaged nephron during later phases of organ repair. Amniotic fluid represents a very suitable source of stem cells for kidney regeneration. AFSCs seem to have great differentiation potential, without the risk of teratoma formation; they also avoid the ethical concerns surrounding embryonic stem cell use, taking into account that amniocentesis is a very safe technique that presents minimal risk to either the mother or the developing fetus. The presence of these preliminary data affirms, nevertheless, that further investigations are still required to confirm the ability of these cells to participate in kidney regeneration, thereby making them beneficial for future therapeutic options.

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26.3.4 Somatic cell nuclear transfer and tissue engineering The application of stem cells for kidney regeneration is very promising and in addition to this particular cell therapy application, there are another two methods that investigators are pursuing to try to restore kidney function for future regenerative medicine applications. One method is somatic cell nuclear transfer and the other involves the tissue engineering approach, using synthetic polymers as scaffolds. Somatic cell nuclear transfer involves the removal of an oocyte nucleus and its replacement with a nucleus derived from a somatic cell obtained from the patient. After this procedure the oocyte is stimulated to undergo multiple divisions using chemicals or electric shock; it develops until it reaches the blastocyst stage and this can eventually be transplanted in utero (reproductive cloning) or embryonic stem cells can be derived and cultured in vitro as cell lines (therapeutic cloning). The first mammal cloned was Dolly54 triggering different scientific groups to investigate in detail the important mechanisms underlying this process. Cloning of other mammals – such as cattle,55 goats,56,57 mice,58 and pigs59–62 – quickly followed using the same techniques. However, the obvious controversy surrounding reproductive cloning63,64 has limited its expansion and therefore investigators have recently focused their attention on therapeutic cloning, mainly because the derivation of embryonic stem cells that can differentiate into various cell lines may provide an alternative source for transplantable cells. An important study by Lanza et al., used therapeutic cloning to produce genetically identical renal tissue in a bovine model.65 The nucleus of a skin fibroblast was microinjected into an enucleated oocyte that was transplanted in utero for 12 weeks and then the cloned renal cells were seeded into a biodegradable scaffold and transplanted in vivo to follow the growth of the tissue-engendered construct. The authors confirmed that the kidney-like organ was capable of secreting urinary fluid, confirming that the implant contained cells capable of filtration, reasorbtion, and secretion. These results were the first demonstration that renal tissue could be created by applying techniques of tissue engineering and therapeutic cloning. It is clear that somatic nuclear transfer technology has many implications for the future, and yet this technology will require more improvement to instill the necessary confidence for its application in real clinical situations. Tissue engineering, combining natural or biodegradable polymers with cells and growth factors, has contributed to the field of kidney regeneration in recent years. The perfect implantable device needs to mimic the main physiological functions of the native kidney and it needs to operate incessantly to remove solutes. Regular dialysis techniques are quite efficient but

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they do not have great adaptability. The target is to design the perfect membrane that has the same filtration capability as the nephron. Humes et al.66 reported the creation of a membrane that has both the pore selectivity and, at the same time, the hydraulic permeability of the native kidney. The creation of better bioartificial hemofilters is important to overcome the problem of loss of filtration due to thrombotic occlusion and protein deposition and that avoid the use of anticoagulant in the extracorporeal units that very often results in bleeding for the patient.67 A few experiments have been conducted where renal cells were cultured in vitro, seeded into a polyglycolic acid polymer scaffold, and subsequently implanted into athimic mice.67 Over time, the formation of nephron-like structures was observed within the polymer. These preliminary results, when improved, could easily be used to produce three-dimensional functional renal structures that can be used in ex vivo or in vivo filtering units. This approach is called ‘cell-based tissue-engineering’ as it refers to the use of scaffolds (natural or synthetic) and cells mixed together to recreate a tissue that mimics the size and functionality of the physiological tissue.

26.4

Conclusions

Regenerative medicine is still in its infancy and requires further insight into the molecular bases of physiological regeneration and injury-induced repair. The twentieth century generated giant advances in scientific and medical knowledge. During the first decade of the twentieth century, the idea of stem cells was introduced by Alexander Maximov as part of his theory of hemotopoiesis.68 The twenty-first century has begun with great excitement and expectations in the field of regenerative medicine. There is new hope that by means of stem cells and innovative technologies some challenging problems in medicine, in particular end stage organ failure, may be overcome. Stem cells originating from both endogenous and exogenous sources are under intense investigation by various research groups all around the world, researching their possible role in renal tissue regeneration in acute and chronic injury models. Embryonic stem cells are the most pluripotential cells among the existing stem cell populations. Their powerful differentiation potential has been used to generate cells of all the three germ layers, including the renal phenotype. However, due to their high tumorigenicity these cells are limited in their effectiveness. For this reason, BM-derived stem cells have been under intense investigation, because they represent a more suitable source of stem cells for renal tissue repair. According to a review of the literature BM-derived stem cells show great plasticity in differentiating into adult cell types of the nephron, more pro-

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foundly into tubular epithelial cells. Whether they are able to ameliorate renal damage in cases of ARF remains to be elucidated. As with BMderived MSCs, AFSCs have the pluripotential to integrate and differentiate into renal structures, as shown by an injection of AFSCs into embryonic kidneys in an in vitro mouse model.52 Whether AFSCs will have a similar effect when applied in an in vivo model remains to be shown. In addition, results seem to indicate that endogenous repair of nephrons due to homing in of progenitor cells within the organ may offer more overall benefit to the patient in the future than exogenous injection of MSCs. There seems to be evidence to support the hypothesis that endogenous epithelial cells and perhaps other progenitors have a key role in the immediate response to damage and repair of renal tubular structures; while perhaps exogenous stem cells, or other sources of MSCs, are mainly responsible for the acute restoration of kidney function by involving secondary mechanisms that regulate or are regulated under an immune cascade. Perhaps both are necessary to achieve the desired effect. In conclusion, we can affirm that different cell- and organ-based approaches using stem cells are being investigated for the purposes of kidney regeneration or kidney function restoration. Further investigations are necessary in order to be able to find the right stem cell populations to apply to cure end stage disease or to provide viable therapeutic options for regenerative medicine applications.

26.5

References

1 perin l, giuliani s, sedrakyan s, da sacco s, de filippo re. Stem cell and regenerative science applications in the development of bioengineering of renal tissue. Pediatr Res. 2008 May;63(5):467–71. 2 thadhani r, pascual m, bonventre jv. Acute renal failure. N Engl J Med. 1996;334:1448–60. 3 kelly kj, molitoris ba. Acute renal failure in the new millennium: Time to consider combination therapy. Semin Nephrol. 2000 Jan;20(1):4–19. 4 grino jm. BN 52021: A platelet activating factor antagonist for preventing posttransplant renal failure. A double-blind, randomized study. The BN 52021 study group in renal transplantation. Ann Intern Med. 1994 Sep 1;121(5):345–7. 5 haug ce, colvin rb, delmonico fl, auchincloss h, jr, tolkoff-rubin n, preffer fi, et al. A phase I trial of immunosuppression with anti-ICAM-1 (CD54) mAb in renal allograft recipients. Transplantation. 1993 Apr;55(4):766–72; discussion 772–3. 6 laflamme ma, murry ce. Regenerating the heart. Nat Biotechnol. 2005 Jul;23(7):845–56. 7 lindvall o, kokaia z. Stem cells for the treatment of neurological disorders. Nature. 2006 Jun 29;441(7097):1094–6. 8 aejaz hm, aleem ak, parveen n, khaja mn, narusu ml, habibullah cm. Stem cell therapy – present status. Transplant Proc. 2007 Apr;39(3):694–9.

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9 clarke dl, johansson cb, wilbertz j, veress b, nilsson e, karlstrom h, et al. Generalized potential of adult neural stem cells. Science. 2000 Jun 2;288(5471):1660–3. 10 blau hm, brazelton tr, weimann jm. The evolving concept of a stem cell: Entity or function? Cell. 2001 105: 829–41. 11 al-awqati q, oliver ja. The kidney papilla is a stem cells niche. Stem Cell Rev. 2006;2(3):181–4. 12 barile l, messina e, giacomello a, marban e. Endogenous cardiac stem cells. Prog Cardiovasc Dis. 2007 Jul–Aug;50(1):31–48. 13 dorrell c, grompe m. Liver repair by intra- and extrahepatic progenitors. Stem Cell Rev. 2005;1(1):61–4. 14 kim cf, jackson el, woolfenden ae, lawrence s, babar i, vogel s, et al. Identification of bronchioalveolar stem cells in normal lung and lung cancer. Cell. 2005 Jun 17;121(6):823–35. 15 poulsom r, forbes sj, hodivala-dilke k, ryan e, wyles s, navaratnarasah s, et al. Bone marrow contributes to renal parenchymal turnover and regeneration. J Pathol. 2001 Sep;195(2):229–35. 16 lin f, cordes k, li l, hood l, couser wg, shankland sj, igarashi p. Hematopoietic stem cells contribute to the regeneration of renal tubules after renal ischemia-reperfusion injury in mice. J Am Soc Nephrol. 2003;14:1188–99. 17 gupta s, verfaillie c, chmielewski d, kim y, rosenberg me. A role for extrarenal cells in the regeneration following acute renal failure. Kidney Int. 2002 Oct;62(4):1285–90. 18 bussolati b, bruno s, grange c, et al. Isolation of renal progenitor cells from adult human kidney. Am J Path. 2005;166(2):545–55. 19 lin f, moran a, igarashi p. Intrarenal cells, not bone marrow-derived cells, are the major source for regeneration in postischemic kidney. J Clin Invest. 2005 Jul;115(7):1756–64. 20 odorico js, kaufman ds, thomson ja. Multilineage differentiation from human embryonic stem cell lines. Stem Cells. 2001;19(3):193–204. 21 humphreys bd, valerius mt, kobayashi a, mugford jw, soeung s, duffield js, mcmahon ap, bonventre jv. Intrinsic epiethelial cells repair the kidney after injury. Cell Stem Cell. 2008;2:284–91. 22 duffield js, park km, hsiao ll, kelley vr, scadden dt, ichimura t, et al. Restoration of tubular epithelial cells during repair of the postischemic kidney occurs independently of bone marrow-derived stem cells. J Clin Invest. 2005 Jul;115(7):1743–55. 23 fang tc, alison mr, cook ht, jeffery r, wright na, poulsom r. Proliferation of bone marrow-derived cells contributes to regeneration after folic acid-induced acute tubular injury. J Am Soc Nephrol. 2005;16:1723–32. 24 held pk, al-dhalimy m, willenbring h, akkari y, jiang s, torimaru y, et al. In vivo genetic selection of renal proximal tubules. Mol Ther. 2006;13:49–58. 25 wang x, willenbring h, akkari y, torimaru y, foster m, al-dhalimy m, et al. Cell fusion is the principal source of bone-marrow-derived hepatocytes. Nature. 2003 Apr 24;422(6934):897–901. 26 weimann jm, johansson cb, trejo a, blau hm. Stable reprogrammed heterokaryons form spontaneously in purkinje neurons after bone marrow transplant. Nat Cell Biol. 2003 Nov;5(11):959–66.

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27 labarge ma, blau hm. Biological progression from adult bone marrow to mononucleate muscle stem cell to multinucleate muscle fiber in response to injury. Cell. 2002 Nov 15;111(4):589–601. 28 ianus a, holz gg, theise nd, hussain ma. In vivo derivation of glucosecompetent pancreatic endocrine cells from bone marrow without evidence of cell fusion. J Clin Invest. 2003 Mar;111(6):843–50. 29 evans mj, kaufman mh. Establishment in culture of pluripotential cells from mouse embryos. Nature. 1981 Jul 9;292(5819):154–6. 30 smith ag, heath jk, donaldson dd, wong gg, moreau j, stahl m, et al. Inhibition of pluripotential embryonic stem cell differentiation by purified polypeptides. Nature. 1988 Dec 15;336(6200):688–90. 31 reubinoff be, pera mf, fong cy, trounson a, bongso a. Embryonic stem cell lines from human blastocysts: Somatic differentiation in vitro. Nat Biotechnol. 2000 Apr;18(4):399–404. 32 kramer j, steinhoff j, klinger m, fricke l, rohwedel j. Cells differentiated from mouse embryonic stem cells via embryoid bodies express renal marker molecules. Differentiation. 2006 Mar;74(2–3):91–104. 33 vigneau c, polgar k, striker g, elliott j, hyink d, weber o, et al. Mouse embryonic stem cell-derived embryoid bodies generate progenitors that integrate long term into renal proximal tubules in vivo. J Am Soc Nephrol. 2007 Jun;18(6):1709–20. 34 abkowitz jl, robinson ae, kale s, long mw, chen j. Mobilization of hematopoietic stem cells during homeostasis and after cytokine exposure. Blood. 2003 Aug 15;102(4):1249–53. 35 iwasaki m, adachi y, minamino k, suzuki y, zhang y, okigaki m, et al. Mobilization of bone marrow cells by G-CSF rescues mice from cisplatin-induced renal failure, and M-CSF enhances the effects of G-CSF. J Am Soc Nephrol. 2005 Mar;16(3):658–66. 36 nishida m, fujimoto s, toiyama k, sato h, hamaoka k. Effect of hematopoietic cytokines on renal function in cisplatin-induced ARF in mice. Biochem Biophys Res Commun. 2004 Nov 5;324(1):341–7. 37 morigi m, imberti b, zoja c, corna d, tomasoni s, abbate m, et al. Mesenchymal stem cells are renotropic, helping to repair the kidney and improve function in acute renal failure. J Am Soc Nephrol. 2004;15:1794–804. 38 herrera mb, bussolati b, bruno s, fonsato v, mauriello-romanazzi g, gamussi g. Mesenchymal stem cells contribute to the renal repair of acute tubular epithelial injury. Int J Mol Med. 2004;14:1035–41. 39 herrera mb, bussolati b, bruno s, morando l, mauriello-romanazzi g, sanavio f, et al. Exogenous mesenchymal stem cells localize to the kidney by means of CD44 following acute tubular injury. Kidney Int. 2007;72:430–41. 40 togel f, hu z, weiss k, isaac j, lange c, westenfelder c. Administered mesenchymal stem cells protect against ischemic acute renal failure through differentiation-independent mechanisms. Am J Physiol Renal Physiol. 2005; 289:F31–F42. 41 lange c, togel f, ittrich h, clayton f, nolte-ernsting c, zander ar, et al. Administered mesenchymal stem cells enhance recovery from ischemia/reperfusion-induced acute renal failure in rats. Kidney Int. 2005 Oct;68(4):1613–7.

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42 yokoo t, fukui a, ohashi t, miyazaki y, utsunomiya y, kawamura t, et al. Xenobiotic kidney organogenesis from human mesenchymal stem cells using a growing rodent embryo. J Am Soc Nephrol. 2006 Apr;17(4):1026–34. 43 ji jf, he bp, dheen st, tay ss. Interactions of chemokines and chemokine receptors mediate the migration of mesenchymal stem cells to the impaired site in the brain after hypoglossal nerve injury. Stem Cells. 2004;22(3):415–27. 44 imberti b, morigi m, tomasoni s, rota c, corna d, longaretti l, et al. Insulin-like growth factor-1 sustains stem cell mediated renal repair. J Am Soc Nephrol. 2007 Nov;18(11):2921–8. 45 jiang y, jahagirdar bn, reinhardt rl, schwartz re, keene cd, ortiz-gonzalez xr, et al. Pluripotency of mesenchymal stem cells derived from adult marrow. Nature. 2002 Jul 4;418(6893):41–9. 46 krause d, cantley lg. Bone marrow plasticity revisited: Protection or differentiation in the kidney tubule? J Clin Invest. 2005 Jul;115(7):1705–8. 47 duffield js, bonventre jv. Kidney tubular epithelium is restored without replacement with bone marrow-derived cells during repair after ischemic injury. Kidney Int. 2005 Nov;68(5):1956–61. 48 gupta s, verfaillie c, chmielewski d, kren s, eidman k, connaire j, et al. Isolation and characterization of kidney-derived stem cells. J Am Soc Nephrol. 2006 Nov;17(11):3028–40. 49 humphreys bd, bonventre jv. Mesenchymal stem cells in acute kidney injury. Annu Rev Med. 2008;59:311–25. 50 togel f, hu z, weiss k, isaac j, lange c, westenfelder c. Administered mesenchymal stem cells protect against ischemic acute renal failure through differentiation-independent mechanisms. Am J Physiol Renal Physiol. 2005 Jul; 289(1):F31–42. 51 de coppi p, bartsch g, jr, siddiqui mm, xu t, santos cc, perin l, et al. Isolation of amniotic stem cell lines with potential for therapy. Nat Biotechnol. 2007 Jan;25(2):100–6. 52 perin l, giuliani s, jin d, sedrakyan s, carraro g, habibian r, et al. Renal differentiation of amniotic fluid stem cells. Cell Prolif. 2007 Dec;40(6):936–48. 53 bonventre jv. Pathophysiology of acute kidney injury: roles of potential inhibitors of inflammation. Contrib Nephrol. 2007; 156:39–46. 54 wilmut i, schnieke ae, mcwhir j, kind aj, campbell kh. Viable offspring derived from fetal and adult mammalian cells. Nature. 1997 Feb 27; 385(6619):810–13. 55 cibelli jb, stice sl, golueke pj, kane jj, jerry j, blackwell c, et al. Cloned transgenic calves produced from nonquiescent fetal fibroblasts. Science. 1998 May 22;280(5367):1256–8. 56 baguisi a, behboodi e, melican dt, pollock js, destrempes mm, cammuso c, et al. Production of goats by somatic cell nuclear transfer. Nat Biotechnol. 1999 May;17(5):456–61. 57 keefer cl, keyston r, lazaris a, bhatia b, begin i, bilodeau as, et al. Production of cloned goats after nuclear transfer using adult somatic cells. Biol Reprod. 2002 Jan;66(1):199–203. 58 wakayama t, perry ac, zuccotti m, johnson kr, yanagimachi r. Full-term development of mice from enucleated oocytes injected with cumulus cell nuclei. Nature. 1998 Jul 23;394(6691):369–74.

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59 betthauser j, forsberg e, augenstein m, childs l, eilertsen k, enos j, et al. Production of cloned pigs from in vitro systems. Nat Biotechnol. 2000 Oct;18(10):1055–9. 60 polejaeva ia, chen sh, vaught td, page rl, mullins j, ball s, et al. Cloned pigs produced by nuclear transfer from adult somatic cells. Nature. 2000 Sep 7;407(6800):86–90. 61 onishi a, iwamoto m, akita t, mikawa s, takeda k, awata t, et al. Pig cloning by microinjection of fetal fibroblast nuclei. Science. 2000 Aug 18; 289(5482):1188–90. 62 de sousa pa, dobrinsky jr, zhu j, archibald al, ainslie a, bosma w, et al. Somatic cell nuclear transfer in the pig: Control of pronuclear formation and integration with improved methods for activation and maintenance of pregnancy. Biol Reprod. 2002 Mar;66(3):642–50. 63 colman a, kind a. Therapeutic cloning: concepts and practicalities. Trends Biotechnol. 2000 May;18(5):192–6. 64 vogelstein b, alberts b, shine k. Genetics. please don’t call it cloning! Science. 2002 Feb 15;295(5558):1237. 65 lanza rp, chung hy, yoo jj, wettstein pj, blackwell c, borson n, et al. Generation of histocompatible tissues using nuclear transplantation. Nat Biotechnol. 2002 Jul;20(7):689–96. 66 humes hd, buffington da, mackay sm, funke aj, weitzel wf. Replacement of renal function in uremic animals with a tissue-engineered kidney. Nat Biotechnol. 1999 May;17(5):451–5. 67 amiel ge, yoo jj, atala a. Renal therapy using tissue-engineered constructs and gene delivery. World J Urol. 2000 Feb;18(1):71–9. 68 maximov aa, Der Lymphozyt als gemeinsame Stammzelle der verschiedened Blutelemente in der embryonalen Entwicklung und im postfetalen Leben der Saugetiere. Folia Haematol (Leipzig). 1909; 8:125–41.

27 Techniques for engineering bladder tissue A. ATA L A, Wake Forest Institute for Regenerative Medicine, USA

Abstract: Congenital disorders, cancer, trauma, infection, inflammation, iatrogenic injuries, or other conditions of the bladder can lead to organ damage or complete loss of function. Most of these situations require eventual reconstructive procedures. These procedures can be performed with native non-urological tissues (skin, gastrointestinal segments, or mucosa), homologous tissues from a donor (cadaver or living donor kidney), heterologous tissues or substances (bovine collagen), or artificial materials (silicone, polyurethane, Teflon). However, these materials often lead to complications after reconstruction. The implanted tissue is sometimes rejected, and often the inherently different functional aspects of the different tissues or materials used in the reconstruction cause a mismatch in the system. As an example, current methods of replacing bladder tissue with gastrointestinal segments can be problematic due to the opposite ways in which these two tissues handle solutes – urological tissue normally excretes material, but gastrointestinal tissue generally absorbs the same materials, and such a mismatch can lead to metabolic complications as well as infection. The replacement of lost or deficient urological tissues with functionally equivalent ones would improve the outcome of reconstructive surgery for the bladder. This may soon be possible with novel tissue engineering techniques. Key words: bladder reconstruction, bladder tissue engineering, stem cells, biomaterials, genitourinary system.

27.1

Introduction

Patients suffering from congenital defects or diseases of the urinary bladder are often treated with reconstructive procedures. Most current techniques involve the use of gastrointestinal segments to repair a portion of the bladder or, in severe cases, create a neobladder. However, the introduction of gastrointestinal tissue into the urogenital tract brings tissues with opposing functions together and creates a mismatch in the system, which can have deleterious effects on both the urinary tract and the patient’s metabolism. In order to find solutions to these problems, physicians and scientists have begun to look to the fields of regenerative medicine and tissue engineering to provide new options for these patients. These fields strive to develop 532

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biological substitutes that can significantly improve the quality of life of the urology patient by eliminating the need for intensive grafting procedures or transplant surgery. Tissue engineering, one of the major components of regenerative medicine, follows the principles of cell transplantation, materials science, and engineering to develop biological substitutes that can restore and maintain normal organ function. Tissue engineering strategies generally fall into two categories: the use of acellular matrices designed to make use of the body’s own infiltrating cells to regenerate damaged tissue, and the use of matrices seeded with cells in the laboratory to produce novel tissues and organs that can be implanted into the patient. In addition, techniques using cells themselves for therapy have been developed. These cells are delivered via injection to the damaged tissue, either with carriers (such as hydrogels) or alone. Acellular tissue matrices are usually prepared by manufacturing artificial scaffolds, or by removing cellular components from donor tissues via mechanical and chemical manipulation to produce collagen-rich matrices.1–4 These matrices slowly degrade after implantation and are replaced by the extracellular matrix (ECM) proteins secreted by the ingrowing cells. The most common way to use cells in tissue engineering is to obtain a small piece of donor tissue and dissociate it into individual cells in the laboratory. These cells are either implanted directly into the host, or are expanded in culture and attached to a support matrix. The cell–matrix construct is then reimplanted into the host. The source of the donor tissue can be heterologous (such as bovine), allogeneic (same species, different individual), or autologous. Ideally, autologous cells are used, because in this case both structural and functional tissue replacement will usually occur with minimal complications.2,5–12 The use of autologous cells may cause a transient inflammatory response, but it avoids rejection and, thus, the deleterious side effects of lifelong immunosuppression can be avoided. However, for many patients with extensive end-stage organ failure, a tissue biopsy may not yield enough normal cells for expansion and transplantation. In other instances, primary autologous cells cannot be expanded from a particular organ, such as the pancreas. In these situations, stem cells are envisioned as an alternative source of cells from which the desired tissue can be derived. Stem cells can be derived from discarded human embryos (human embryonic stem (hES) cells), from fetal tissue, or from adult sources (bone marrow, fat, skin). However, there are ethical issues involved in the use of ES cells and most human applications are currently banned in the United States. Despite this, the field of stem cell biology is advancing rapidly, and cutting-edge techniques such as therapeutic cloning and somatic cell reprogramming circumvent some of the ethical questions and offer a potentially limitless source of these cells for tissue engineering applications. This chapter will review the major components of most tissue engineering tech-

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niques, and will describe how these techniques are being applied to the reconstruction and regeneration of the urinary bladder.

27.2

Cells used in tissue engineering

27.2.1 Native cells In the past, one of the limitations of applying cell-based regenerative medicine techniques to organ replacement was the inherent difficulty of growing certain cell types in large quantities. Even when some organs, such as the liver, have a high regenerative capacity in vivo, cell growth and expansion in vitro can be difficult. By studying the privileged sites for committed precursor cells in these organs, as well as by exploring the conditions that promote differentiation and/or self-renewal of these cells, it has been possible to overcome some of the obstacles that limit cell expansion in vitro. One example is the urothelial cell. Urothelial cells could be grown in the laboratory setting in the past, but only with limited success. Several protocols have been developed over the past two decades that identify the undifferentiated cells in a mixed culture of cells isolated from the urinary tract, and keep them undifferentiated during their growth phase.11,13–16 Using these methods of cell culture, it is now possible to expand a urothelial culture that initially covered a surface area of 1 cm2 to one covering a surface area of 4202 m2 (the equivalent of one football field) within 8 weeks.11 These studies indicated that it should be possible to collect autologous bladder cells from human patients, expand them in culture, and return them to the donor in sufficient quantities for reconstructive purposes.11,14–19 Major advances in cell culture techniques have been made within the past decade, and these techniques may make the use of autologous cells a clinical reality.

27.2.2 Embryonic stem cells In 1981, pluripotent cells were found in the inner cell mass of the human embryo, and the term ‘human embryonic stem cell’ was coined.20 These cells are able to differentiate into all cells of the human body, excluding placental cells (only cells from the morula are totipotent; that is, able to develop into all cells of the human body including the placenta). These cells have great therapeutic potential, but their use is limited by both biological and ethical factors. The political controversy surrounding stem cells began in 1998 with the creation of hES cells derived from discarded embryos. hES cells were isolated from the inner cell mass of a blastocyst (an embryo 5 days postfertilization) using an immunosurgical technique. As previously discussed,

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some cells cannot be expanded ex vivo, and in these cases ES cells could be an ideal resource for tissue engineering because of their fundamental properties: the ability to self-renew indefinitely and the ability to differentiate into cells from all three embryonic germ layers. Skin and neurons have been formed, indicating ectodermal differentiation.21–24 Blood, cardiac cells, cartilage, endothelial cells, and muscle have been formed, indicating mesodermal differentiation.25–27 Finally, pancreatic cells have been formed, indicating endodermal differentiation.28 In addition, as further evidence of their pluripotency, ES cells can form embryoid bodies, which are cell aggregations that exist in culture and contain all three embryonic germ layers, and they can also form teratomas in vivo.29 These cells have demonstrated longevity in culture and can maintain their undifferentiated state for at least 80 passages when grown using current published protocols.30,31 However, in addition to the ethical dilemma surrounding the use of ES cells, their clinical application is also limited because they represent an allogenic resource and thus have the potential to evoke an immune response. New stem cell technologies (such as somatic cell nuclear transfer and reprogramming) promise to overcome this limitation.

27.2.3 Therapeutic cloning (somatic cell nuclear transfer) Somatic cell nuclear transfer (SCNT), or therapeutic cloning, entails the removal of an oocyte nucleus in culture, followed by its replacement with a nucleus derived from a somatic cell obtained from a patient. Activation with chemicals or electricity stimulates cell division up to the blastocyst stage, and ES cells that are genetically identical to the patient can be obtained from the newly generated inner cell mass. At this point in our discussion, it is extremely important to differentiate between the two types of cloning that exist – reproductive cloning and therapeutic cloning. Both involve the insertion of donor DNA into an enucleated oocyte to generate an embryo that has identical genetic material to its DNA source. However, the similarities end there. In reproductive cloning, the embryo is then implanted into the uterus of a pseudopregnant female to produce an infant that is a clone of the donor. A world-famous example of this type of cloning resulted in the birth of a sheep named Dolly in 1997.32 However, there are many ethical concerns surrounding such practices and, as a result, reproductive cloning has been banned in most countries. While therapeutic cloning also produces an embryo that is genetically identical to the donor, this process is used to generate blastocysts that are explanted and grown in culture, rather than in utero. ES cell lines can then be derived from these blastocysts, which are only allowed to grow up to a 100-cell stage. At this time the inner cell mass is isolated and cultured,

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Enucleated oocyte

Nuclear transfer

Embryonic stem cells

Transplantation

Hematopoietic cells

Genetically matched tissue Neurons

Pancreatic islet cells Cardiomyocytes

Hepatocytes Renal cells

27.1 Strategies for therapeutic cloning and tissue engineering.

resulting in ES cells that are genetically identical to the patient. This process is detailed in Fig. 27.1. It has been shown that ES cells derived from the nuclei of fibroblasts, lymphocytes, and olfactory neurons are pluripotetent and can generate live pups after injection into blastocysts. This shows that cells generated by SCNT have the same developmental potential as blastocysts that are fertilized and produced naturally.33–36 In addition, the ES cells generated by SCNT are perfectly matched to the patient’s immune system and no immunosuppressants would be required to prevent rejection should these cells be used in tissue engineering applications. Although ES cells derived from SCNT contain the nuclear genome of the donor cells, mitochondrial DNA (mtDNA) contained in the oocyte could lead to immunogenicity after transplantation. In order to assess the histocompatibilty of tissue generated using SCNT, Lanza et al.37 microinjected the nucleus of a bovine skin fibroblast into an enucleated oocyte.37 Although the blastocyst was implanted (reproductive cloning), the purpose was to generate renal, cardiac, and skeletal muscle cells, which were then harvested, expanded in vitro, and seeded on to biodegradable scaffolds. These scaffolds were then implanted into the donor steer from which the cells were cloned to determine if the cells were histocompatible. Analysis revealed no evidence of a mounting T-cell response to the cloned renal cells, suggesting that rejection will not necessarily occur in the presence of oocytederived mtDNA. This finding represents a step forward in overcoming the histocompatibility problem of stem cell therapy.

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Although promising, SCNT has certain limitations that require further improvement before its clinical application, in addition to the ethical considerations regarding the potential of SCNT-derived embryos to develop into offspring if implanted into a uterus. First and foremost, this technique has not yet been shown to work in humans. The initial failures and fraudulent reports of nuclear transfer in humans reduced excitement in the scientific community and suggested that SCNT may not be ideal for human applications,38–40 although it was recently reported that non-human primate ES cell lines were generated by SCNT of nuclei from adult skin fibroblasts.41,42 In addition, before SCNT-derived ES cells can be used for any type of clinical therapy, careful assessment of the quality of the generated cell lines must be made. For example, some cell lines generated by SCNT have contained chromosomal translocations and it is not known whether these abnormalities originated from the generation of aneuploid embryos, or if they occurred during ES cell isolation and culture. The low ‘success rate’ of SNCT (only 0.7% of the nuclear transfers result in a viable blastocyst) and the inadequate supply of human oocytes further hinder the therapeutic potential of this technique. Still, these studies have renewed the hope that ethically neutral ES cell lines could one day be generated from human cells to produce patient-specific stem cells for use in tissue engineering and regenerative medicine applications.

27.2.4 Reprogrammed somatic cells Recently, exciting reports of the successful transformation of adult cells into pluripotent stem cells through a type of genetic ‘reprogramming’ have been published. Reprogramming is a technique that involves dedifferentiation of adult somatic cells to produce patient-specific pluripotent stem cells, thus eliminating the need to create embryos to obtain stem cells. Stem cells generated by reprogramming would be genetically identical to the somatic cells used to create them (and, thus, the patient who donated those cells) and would not be rejected. Takahashi and Yamanaka were the first to discover that mouse embryonic fibroblasts (MEFs) and adult mouse fibroblasts could be reprogrammed into an ‘induced pluripotent state (iPS)’.43 These iPS cells possessed the immortal growth characteristics of self-renewing ES cells, expressed genes specific for ES cells, and generated embryoid bodies in vitro and teratomas in vivo. When iPS cells were injected into mouse blastocysts, which were allowed to divide, they contributed to a variety of cell types during development. However, although iPS cells selected in this way were pluripotent, they were not identical to ES cells. Unlike ES cells, chimeras made from iPS cells did not result in full-term pregnancies. Gene expression profiles of the iPS

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cells showed that they possessed a distinct gene expression signature that was different from that of ES cells. In addition, the epigenetic state of the iPS cells was somewhere between that found in somatic cells and that found in ES cells, suggesting that the reprogramming was incomplete. These results were improved significantly by Wernig et al. in July 2007.44 In their reprogrammed cells, the DNA methylation pattern, gene expression profiles, and chromatin state were similar to those of ES cells. Teratomas induced by these cells contained differentiated cell types representing all three embryonic germ layers. Most importantly, the reprogrammed cells from this experiment were able to form viable chimeras and contribute to the germ line like true ES cells, suggesting that these iPS cells were completely reprogrammed. It has recently been shown that reprogramming of human cells is possible.45,46 Yamanaka generated human iPS cells that are similar to hES cells in terms of morphology, proliferation, gene expression, surface markers, and teratoma formation. Thompson’s group showed that retroviral transduction of the stem cell markers OCT4, SOX2, NANOG, and LIN28 could generate pluripotent stem cells. However, in both studies, the human iPS cells were similar but not identical to hES cells, and there is always cause for concern when viruses are introduced into cell lines, since the long-term effects of the presence of a retroviral genome in engineered cells and tissues are unclear at present. Although reprogramming is an exciting phenomenon, our limited understanding of the mechanism underlying it currently limits the clinical applicability of the technique, but the future potential of reprogramming is quite intriguing.

27.2.5 Placental and amniotic fluid stem cells Recently, it has been shown that pluripotent cells may be derived from the amniotic fluid and placenta. Both amniotic fluid and placenta are known to contain multiple partially differentiated cell types derived from the developing fetus. In the past decade, stem cell populations have been isolated from these sources. Called ‘amniotic fluid and placental stem cells (AFPSCs)’, they express embryonic and adult stem cell markers.47 The undifferentiated stem cells expand extensively without a feeder cell layer and double every 36 hours. Unlike hES cells, the AFPSCs do not form tumors in vivo. Lines maintained for over 250 population doublings retained long telomeres and a normal complement of chromosomes. AFS cells are broadly multipotent, and human lines can be induced to differentiate into cell types representing each embryonic germ layer, including cells of adipogenic, osteogenic, myogenic, endothelial, neuronal, hepatic, and renal lineages. Examples of differentiated cells derived from AFS cells and displaying specialized functions include neuronal cells that secrete the neurotransmitter l-glutamate or

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express G-protein-gated inwardly rectifying potassium (GIRK) channels, hepatic lineage cells that produce urea, and osteogenic cells that can form tissue engineered bone. In this respect, AFPSCs meet a commonly accepted criterion for pluripotent stem cells, without implying that they can generate every adult tissue. The cells can be obtained either from amniocentesis or chorionic villous sampling in the developing fetus, or from the placenta at the time of birth. They could be preserved for self use, and used without rejection, or they could be banked. A bank of 100 000 specimens could potentially supply 99% of the US population with a perfect genetic match for transplantation. Such a bank may be easier to create than with other cell sources, since there are approximately 4.5 million births per year in the USA.47

27.3

Biomaterials used in tissue engineering

In the most common tissue engineering procedures, isolated cells are seeded on to a scaffold composed of an appropriate biomaterial. These biomaterials replicate the biological and mechanical function of the native ECM found in tissues in the body by serving as an artificial ECM. Biomaterials provide a three-dimensional space for the cells to develop into new tissues with appropriate structure and function. They can also allow delivery of appropriate bioactive factors (e.g. cell adhesion peptides, growth factors) to the developing tissue48 to help regulate cellular function. As the majority of mammalian cell types are anchorage-dependent and will die if no cell adhesion substrate is available, biomaterials provide this substrate that can deliver cells to specific sites in the body with high loading efficiency. Biomaterials can also provide mechanical support against in vivo forces so that the predefined three-dimensional structure of the engineered implant is maintained during tissue development. The ideal biomaterial should be biodegradable, bioresorbable, and support the replacement of normal tissue without inducing inflammation. Incompatible materials are destined for an inflammatory or foreign-body response that eventually leads to rejection and/or necrosis. Degradation products, if produced, should be removed from the body via normal metabolic pathways at an adequate rate so that the concentration of these degradation products in the tissue remains at a tolerable level.49 The biomaterial should also provide an environment in which appropriate regulation of cell behavior (adhesion, proliferation, migration, and differentiation) can occur. Cell behavior in the newly formed tissue has been shown to be regulated by multiple interactions of the cells with their microenvironment, including interactions with cell adhesion ligands50 and with soluble growth factors. Since biomaterials provide temporary mechanical support while the cells undergo spatial reorganization into tissue, the

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properly chosen biomaterial should allow the engineered tissue to maintain sufficient mechanical integrity to support itself in early development, while in late development it should begin to degrade so that it does not hinder further tissue growth.48 Generally, three classes of biomaterials have been utilized for engineering tissues: naturally derived materials (e.g. collagen and alginate), acellular tissue matrices (e.g. bladder submucosa and small intestinal submucosa), and synthetic polymers such as polyglycolic acid (PGA), polylactic acid (PLA), and poly(lactic-co-glycolic acid) (PLGA). These classes of biomaterials have been tested in respect to their biocompatibility.51,52 Naturally derived materials and acellular tissue matrices have the potential advantage of biological recognition. However, synthetic polymers can be produced reproducibly on a large scale with controlled properties such as strength, degradation rate, and microstructure – which would aid in the preparation of easily used, ‘off-the-shelf’ scaffold material.

27.3.1 Naturally derived materials Collagen is the most abundant and ubiquitous structural protein in the body, and may be readily purified from both animal and human tissues with an enzyme treatment and salt/acid extraction.53 Collagen implants, under normal conditions, degrade through a process involving phagocytosis of collagen fibrils by fibroblasts.54 This is followed by sequential attack by lysosomal enzymes including cathepsins B1 and D. Under inflammatory conditions, the implants can be rapidly degraded largely by matrix metalloproteins (MMPs) and collagenases.54 However, the in vivo resorption rate of a collagen implant can be regulated by controlling the density of the implant and the extent of intermolecular crosslinking – the lower the density, the greater the space between collagen fibers and the larger the pores for cell infiltration, leading to a higher rate of implant degradation. Collagen contains cell adhesion domain sequences (e.g. RGD (arginine–glycine–aspartic acid)) that may assist in retaining the phenotype and activity of many types of cells, including fibroblasts55 and chondrocytes.56 Alginate, a polysaccharide isolated from seaweed, has been used as an injectable cell delivery vehicle57 and a cell immobilization matrix58 owing to its gentle gelling properties in the presence of divalent ions such as calcium. Alginate is relatively biocompatible and is approved by the US Food and Drug Administration (FDA) for human use as wound dressing material. Alginate is a family of copolymers of d-mannuronate and l-guluronate. The physical and mechanical properties of alginate gel are strongly correlated with the proportion and length of polygluronic block in the alginate chains.57

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27.3.2 Acellular tissue matrices Acellular tissue matrices are collagen-rich matrices prepared by removing cellular components from tissues. The matrices are often prepared by mechanical and chemical manipulation of a segment of tissue.1–4 These matrices slowly degrade upon implantation, and are replaced and remodeled by ECM proteins synthesized and secreted by transplanted cells or in growing cells.

27.3.4 Synthetic polymers Polyesters of naturally occurring α-hydroxy acids – including PGA, PLA, and PLGA – are widely used in tissue engineering. These polymers are FDAapproved for a variety of applications, including sutures.59 The ester bonds in these polymers are hydrolytically labile, and they degrade by non-enzymatic hydrolysis. The degradation products of PGA, PLA, and PLGA are nontoxic natural metabolites and are eventually eliminated from the body in the form of carbon dioxide and water.59 The degradation rate of these polymers can be tailored to the application by altering crystallinity, initial molecular weight, and the copolymer ratio of lactic to glycolic acid. Generally, the optimal degradation time ranges from several weeks to several years. Since these polymers are thermoplastics, they can be easily formed into a threedimensional scaffold with a desired microstructure, gross shape, and dimension by various techniques – including molding, extrusion, solvent casting,60 phase separation techniques, and gas foaming techniques.61 Many applications in tissue engineering often require a scaffold with high porosity and ratio of surface area to volume. Other biodegradable synthetic polymers, including poly(anhydrides) and poly(ortho-esters), can also be used to fabricate scaffolds for tissue engineering with controlled properties.62

27.4

Bladder repair and replacement: current and future technologies

Currently, gastrointestinal segments are commonly used for bladder replacement or repair. However, gastrointestinal tissues are designed to absorb solutes that urinary tissue excretes, and due to this difference in function, multiple complications may ensue, such as infection, metabolic disturbances, urolithiasis, perforation, increased mucus production, and malignancy.63–65 Because of the problems encountered with the use of gastrointestinal segments, numerous investigators have attempted alternative reconstructive procedures for bladder replacement or repair. The use of tissue expansion, seromuscular grafts, matrices for tissue regeneration, and tissue engineering with cell transplantation have been investigated.

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27.4.1 Tissue expansion for bladder augmentation A system of progressive dilation for ureters and bladders has been proposed as a method of bladder augmentation but has not yet been attempted clinically. Augmentation cystoplasty performed with dilated ureteral segments in animals has resulted in an increased bladder capacity ranging from 190% to 380%.66,67 A system for the progressive expansion of native bladder tissue has also been used for augmenting bladder volumes in animals. Within 30 days after progressive dilation, the neoreservoir volume was expanded at least 10-fold. Urodynamic studies showed normal compliance in all animals and microscopic examination of the expanded neoreservoir tissue showed a normal histology. A series of immunocytochemical studies demonstrated that the dilated bladder tissue maintained normal phenotypic characteristics.67

27.4.2 Seromuscular grafts and de-epithelialized bowel segments Seromuscular grafts and de-epithelialized bowel segments, either alone or over a native urothelium, have also been attempted.68–73 Keeping the urothelium intact avoids the complications associated with the use of bowel in continuity with the urinary tract.69,70 An example of this strategy is the combination of the techniques of autoaugmentation with those of enterocystoplasty. An autoaugmentation is performed and the diverticulum is covered with a demucosalized gastric or intestinal segment.42

27.4.3 Matrices for bladder regeneration Non-seeded allogeneic acellular matrices have served as scaffolds for the ingrowth of host bladder wall components. The matrices are prepared by mechanically and chemically removing all cellular components from bladder tissue.2,3,74,75 The matrices serve as vehicles for partial bladder regeneration, and relevant antigenicity is not evident. One example is small intestinal submucosa (SIS), a biodegradable, acellular, xenogeneic collagen-based tissue matrix. SIS was first used in the early 1980s as an acellular matrix for tissue replacement in the vascular field. It has been shown to promote regeneration of a variety of host tissues, including blood vessels and ligaments.76 Animal studies have shown that the non-seeded SIS matrix used for bladder augmentation is able to regenerate in vivo.77,78 In multiple studies using various materials as non-seeded grafts for cystoplasty, the urothelial layer was able to regenerate normally, but the muscle layer, although present, was not fully developed.2,74,75,78 Often the grafts contracted to 60%–70% of their original size with little increase in bladder

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capacity or compliance.79,80 Studies involving acellular matrices that may provide the necessary environment to promote cell migration, growth, and differentiation are being conducted. Recently, bladder regeneration has been shown to be more reliable when the SIS was derived from the distal ileum.77 With continued research in this area, these matrices may have a clinical role in bladder replacement in the future.

27.4.4 Bladder replacement using tissue engineering Cell-seeded allogeneic acellular bladder matrices have been used for bladder augmentation in dogs. A group of experimental dogs underwent a trigone-sparing cystectomy and were randomly assigned to one of three groups. One group underwent closure of the trigone without a reconstructive procedure, another underwent reconstruction with a non-seeded bladder-shaped biodegradable scaffold, and the last underwent reconstruction using a bladder-shaped biodegradable scaffold that was seeded with autologous urothelial and smooth muscle cells.12 The cystectomy-only and non-seeded controls maintained average capacities of 22% and 46% of preoperative values, respectively. However, an average bladder capacity of 95% of the original precystectomy volume was achieved in the cell-seeded tissue engineered bladder replacements (Fig. 27.2). The subtotal cystectomy reservoirs that were not reconstructed and the polymer-only reconstructed bladders showed a marked decrease in bladder compliance (10% and 42% total compliance, respectively). The compliance of the cell-seeded tissue engineered bladders was almost no different from preoperative values (106%). Histologically, the non-seeded scaffold bladders presented a pattern of normal urothelial cells with a thickened fibrotic submucosa and a thin layer of muscle fibers. The retrieved tissue engineered bladders showed a normal cellular organization, consisting of a trilayer of urothelium, submucosa, and muscle.12 A clinical experience involving engineered bladder tissue for cystoplasty reconstruction was conducted starting in 1999. A small pilot study of seven patients was reported, using a collagen scaffold seeded with cells either with or without omentum coverage, or a combined PGA–collagen scaffold seeded with cells and omental coverage (see Fig. 27.3). The patients reconstructed with the engineered bladder tissue created with the PGA–collagen cell-seeded scaffolds showed increased compliance, decreased end-filling pressures, increased capacities, and longer dry periods (Fig. 27.4).81 Although the experience is promising in terms of showing that engineered tissues can be implanted safely, it is just a start in terms of accomplishing the goal of engineering fully functional bladders. Further experimental and clinical work are being conducted.

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(a)

(b)

(c)

Subtotal cystectomy only

Polymer only implants

Tissue engineered neobladder

11 M

11 M

11 M

27.2 Gross specimens and cystograms at 11 months of the cystectomy-only (a), non-seeded controls (b), and cell-seeded tissue engineered bladder replacements (c) from experimental dogs. The cystectomy-only bladder had a capacity of 22% of the preoperative value and a decrease in bladder compliance to 10% of the preoperative value. The non-seeded controls showed significant scarring with a capacity of 46% of the preoperative value and a decrease in bladder compliance to 42% of the preoperative value. An average bladder capacity of 95% of the original precystectomy volume was achieved in the cell-seeded tissue engineered bladder replacements and the compliance showed almost no difference from preoperative values that were measured when the native bladder was present (106%).

(a)

(b)

(c)

27.3 Construction of engineered human bladder. (a) Engineered bladder anastamosed to native bladder with running 4-0 polyglocolic sutures. (b) Implant covered with fibrin glue and omentum.

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(a) Pves 30 cmH2O/div t1 t3 t2

t4 t6 t5 t7

t8 t9 t11 t13 t10 t12

3 min/div

(b)

Pves 30 cmH2O/div t1

5 min/div t2 t4 t5 t3

27.4 Cystograms and urodynamic studies of a patient before and after implantation of the tissue engineered bladder. (a) Preoperative results indicate an irregular bladder in the cystogram and abnormal bladder pressures (Pves) as the bladder is filled via urodynamic study. (b) Postoperatively, findings are significantly improved.

27.5

Summary and conclusions

Regenerative medicine efforts are currently underway experimentally for virtually every type of tissue and organ within the human body. These tissues are at different stages of development, with many still in the discovery stage. However, some tissues, particularly those of the urinary tract, are in clinical trials or are already in use in urological surgical procedures. Recent progress suggests that engineered urological tissues may have an expanded clinical applicability in the future and may represent a viable therapeutic option for those who would benefit from bladder replacement or repair.

27.6

References

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62 peppas, n. a., langer, r.: New challenges in biomaterials. [see comment]. Science, 263: 1715, 1994. 63 mcdougal, w. s.: Metabolic complications of urinary intestinal diversion. J Urol, 147: 1199, 1992. 64 kaefer, m., hendren, w. h., bauer, s. b. et al.: Reservoir calculi: a comparison of reservoirs constructed from stomach and other enteric segments. [see comment]. J Urol, 160: 2187, 1998. 65 kaefer, m., tobin, m. s., hendren, w. h. et al.: Continent urinary diversion: the Children’s Hospital experience. J Urol, 157: 1394, 1997. 66 lailas, n. g., cilento, b., atala, a.: Progressive ureteral dilation for subsequent ureterocystoplasty. J Urol, 156: 1151, 1996. 67 satar, n., yoo, j. j., atala, a.: Progressive dilation for bladder tissue expansion. J Urol, 162: 829, 1999. 68 blandy, j. p.: Ileal pouch with transitional epithelium and anal sphincter as a continent urinary reservoir. J Urol, 86: 749, 1961. 69 blandy, j. p.: The feasibility of preparing an ideal substitute for the urinary bladder. Ann R Coll Surg Engl, 35: 287, 1964. 70 harada, n., yano, h., ohkawa, t. et al.: New surgical treatment of bladder tumours: mucosal denudation of the bladder. Br J Urol, 37: 545, 1965. 71 oesch, i.: Neourothelium in bladder augmentation. An experimental study in rats. Eur Urol, 14: 328, 1988. 72 salle, j. l., fraga, j. c., lucib, a. et al.: Seromuscular enterocystoplasty in dogs. J Urol, 144: 454, 1990. 73 cheng, e., rento, r., grayhack, j. t. et al.: Reversed seromuscular flaps in the urinary tract in dogs. J Urol, 152: 2252, 1994. 74 probst, m., dahiya, r., carrier, s. et al.: Reproduction of functional smooth muscle tissue and partial bladder replacement. Br J Urol, 79: 505, 1997. 75 sutherland, r. s., baskin, l. s., hayward, s. w. et al.: Regeneration of bladder urothelium, smooth muscle, blood vessels and nerves into an acellular tissue matrix. J Urol, 156: 571, 1996. 76 badylak, s. f., lantz, g. c., coffey, a. et al.: Small intestinal submucosa as a large diameter vascular graft in the dog. J Surg Res, 47: 74, 1989. 77 kropp, b. p., cheng, e. y., lin, h. k. et al.: Reliable and reproducible bladder regeneration using unseeded distal small intestinal submucosa. J Urol, 172: 1710, 2004. 78 kropp, b. p., rippy, m. k., badylak, s. f. et al.: Regenerative urinary bladder augmentation using small intestinal submucosa: urodynamic and histopathologic assessment in long-term canine bladder augmentations. J Urol, 155: 2098, 1996. 79 portis, a. j., elbahnasy, a. m., shalhav, a. l. et al.: Laparoscopic augmentation cystoplasty with different biodegradable grafts in an animal model. J Urol, 164: 1405, 2000. 80 landman, j., olweny, e., sundaram, c. p. et al.: Laparoscopic mid sagittal hemicystectomy and bladder reconstruction with small intestinal submucosa and reimplantation of ureter into small intestinal submucosa: 1-year followup. J Urol, 171: 2450, 2004. 81 atala, a., bauer, s. b., soker, s. et al.: Tissue-engineered autologous bladders for patients needing cystoplasty. Lancet, 367: 1241, 2006.

Index

abdominal leak point pressure, 409 Accuflex stent, 109, 113 acellular bladder matrix, 478 acellular dermal matrix, 457 acetohydroxamic acid, 73, 178–9 activin A, 381 adiponectin, 398 adipose progenitor cells applications in urology, 395–412, 404–11 bladder, 406–7 erectile dysfunction and male infertility, 410–11 stress urinary incontinence treatment, 408–10 urethra, 407–8 differentiation capacity of adiposederived stem cells, 402–4 future trends, 411–12 isolation procedures, 397–9 molecular characterisation, 399–401 molecular immunophenotype, 400 nomenclature and origin, 397 adipose stroma vascular cell. see adipose progenitor cells adipose tissue classification bone marrow, 398 brown, 398 mammary, 398 mechanical, 398 white, 398 adipose tissue-derived stromal cells see adipose progenitor cells

550

adipose-derived stem cells, 397 see also adipose progenitor cells vs mesenchymal stem cells, 401 adult stem cells, 423 advanced catheter system, 181 Aequorea victoria, 350 albumin, 164, 390 alginate, 494, 540 as fundamental natural biomaterial, 257–9 applications, 258–9 vs alternative materials, 257–8 alkaline phosphatase, 385 alkoxy condensation reaction system, 192 alkoxysilane, 192 ‘four-armed’ crosslinker, 197 AlloDerm, 273, 457 α–β-methylene-ATP, in endogenous bladder regeneration, 324, 326 alpha-fetoprotein, 390 α-1-microglobulin, 164 α-minimal essential medium, 495, 496 alpha-myosin heavy chain, 381 alpha-smooth muscle actin, 402, 403, 405 for smooth muscle cells characterisation, 339 AMD3100, 323 American Medical Systems, 227, 231, 232, 234 amniocentesis, 378–9 amniotic epithelial cells, human, 381–2

Index amniotic fluid, 371 amniocentesis, 378–9 derived stem cells cell cloning and retroviral marking, 384 cell preparation and culture methods, 382–3 isolation and characterisation, 383–4 multilineage induction, 384–90 differentiated cells, 379 mesenchymal stem cells, 379–82 multilineage induction adipogenic differentiation, 384–5 endothelial differentiation, 387–8 glutamic acid secretion induced by potassium ions, 389 hepatic differentiation, 389–90 myogenic differentiation, 386–7 neurogenic differentiation, 388–9 osteogenic differentiation, 385–6 and placental stem cells, 524, 538–9 as source for urological regenerative medicine, 378–90 ‘amniotic fluid stem cells,’ 382 amniotic membrane, 458–9 collagen types basal lamina collagen type IV, 458 fibrillar collagen types I and III, 458 growth factors basic fibroblast growth factor, 458 epidermal growth factor, 458 hepatocyte growth factor, 458 keratinocyte growth factor, 458 amniotic mesenchymal cells, human, 381 animal model commonly used for ureteral stents, 45–54 dog, 50–2 rabbit, 47–50 rat, 45, 47 swine, 52–4 in vivo models for ureteral stents, 42–55

551

comparative normal urine biochemistry, 46 history, 42–3 vs in vitro models, 43–4 antimicrobials for biofilm prevention and treatment, 26–8 hydrogels, 27 silver, 26–7 triclosan, 27–8 anti-pancytokeratins, 338 artificial biomaterials for urological tissue engineering, 243–52 Assing FESEM Supra 25, 138 atomic absorption spectroscopy, 70, 71, 72 attenuated total reflectance, 138, 143, 151 autoaugmentation, 542 autologous cartilage tissue, 477 autologous cells cell tracking technology, 348, 350–1 fully differentiated cells for urological reconstruction, 337–42 cells of non-urological origin, 340–1 multilayered urological sheet, 337–8 smooth muscle cells characterisation, 339–40 urological smooth muscle cells, 339 urothelial cells, 337 urothelial cells characterisation, 338 uses in urinary tract reconstruction, 341–2 methods of tracking grafted cells and cellular markers, 349 potential use for urological tissue engineering and cell therapy, 336 sources for urological applications, 334–51

552

Index

stem/progenitor cells for urological reconstruction, 342–8 cells from non-urological tissues, 344–5 cells from urinary tract, 342–4 for endoscopic therapies, 346–8 methods of inducing differentiation, 345–6 autologous ear chondrocytes, 423 Barnes stent, 214 benign prostate enlargement, 219 bioabsorption and biodegradation, 215 definition, 215 bioactive silicone, 136 biodegradation and bioabsorption, 215 definition, 215 biofilms bacterial formation and development, 6 characteristics, 135 conditioning film, 7 definition, 61 description and purpose, 4–5 formation and catheter design, 157–85 formation and structure, 5–12 communication, 11–12 effects of hydrodynamics, 9–10 future trends, 30–1 in general medicine, 12–14 Klebsiella pneumoniae biofilm on ureteral stent, 8 models for assessment of formation on urological materials, 59–78 prevention and treatment strategies antimicrobials, 26–8 biofilm-disrupting agents, 28 novel antifouling coatings, 29 novel stent designs, 29–30 resistance, 23–5 antibiotics, 24–5 host factors, 23–4 shedding and migration, 22–3

stages of formation, 135 ureteral stent biofilms involving major bacterial species infection, 19 in urology, 3–31, 14–18, 20–2 associated urinary infections, 15–18, 20 female urogenital tract, 20–1 IBCs caused by uropathogenic Escherichia coli, 21–2 upper and lower urinary tract hydrodynamics, 14–15 bioflex, for cylinders and reservoir bladders, 235 ‘biological signals,’ 330 biomaterials alginate polysaccharide, 257 artificial, for urological tissue engineering, 243–52 future trends, 251–2 general considerations, 245–6 history, 244–5 specific considerations, 246–7 collagen microstructure, 260 collagen-based extracellular matrices, 261–3, 265–74 acellular dermal matrix, 273–4 bladder extracellular matrix, 268–73 porcine small intestinal submucosa, 262–3, 265–7 fundamental, 257–61 alginate, 257–9 collagen, 259–61 natural, for urological tissue engineering, 255–75 acetylated bovine pericardium, 256–7 future trends, 274–5 historical application, 256–7 lyophilised human dura, 256 placental membranes, 257 SIS physical structure variation, 264 smart, 247–51 collagen, 248–9 elastin, 249

Index fibronectin, 249 glycosaminoglycans, 250 growth factors and cytokines, 250–1 ‘smart bladder construct,’ 252 synthetic scaffolds, 245–7 use of nanomaterials, 283–6 biomimetic silicone, 136 ‘bioreactor,’ 329 bladder, 301, 304 engineered bladder development, 328–30 biomaterials and scaffolds, 329–30 cellular source, 328–9 ‘smart’ scaffolds development, 330 evaluation of engineered or regenerating tissues in vitro, 304–10 biochemical evaluation, 309–10 contractile and relaxation responses, 304–7 intracellular messenger molecules real-time fluorescence imaging, 307–8 trans-membrane ionic currents, 308–9 implantation in preclinical studies, 330–1 preliminary clinical experience with neobladders, 331 regeneration, and regenerative pharmacology, 322–32 regional variation in response to electrical field stimulation, 326 repair and replacement, 541–3 bladder replacement using tissue engineering, 543 matrices for bladder regeneration, 542–3 seromuscular grafts and de-epithelialised bowel segments, 542 tissue expansion for bladder augmentation, 542 tissue engineering and regeneration, 310, 312–14

553

evaluating de novo bladder regeneration, 313–14 evaluating tissue-engineered bladders, 312–13 functional performance evaluation matrix, 305 methods of studying tissueengineered bladder constructs, 311 outward potassium currents, 309 pharmacological studies conducted in organ bath apparatus, 306 quantitative changes in fura-2 fluorescence in smooth muscle cells, 308 tissue engineering literature review, 302–3 tissue engineering techniques, 532–45 tissue-engineered implant performance assessment, 299–315 future trends, 314–15 bladder acellular matrix allograft, 271 bladder acellular matrix graft, 460–4 acellular matrices vs cell-seeded matrices, 461 in tissue repair, 460 in treatment of urethral pathology, 462 vs buccal mucosa grafts, 462–3 bladder extracellular matrix, 268–73 effect of cell seeding on regenerative properties, 272 fundamentals, 268–9 pre-clinical studies, 270–3 vs SIS, 268–9 bone marrow stroma, 427 bone marrow-derived stem cells, 344–5 endogenous, 520 exogenous, 522–4 hematopoietic stem cells, 522–3 mesenchymal stem cells, 523–4 bovine collagen, 423 serum albumin, 496

554

Index

bromodeoxyuridine, 508 brushite, 71, 135, 147, 153, 154 burn injuries, 13 cadaveric fascia lata, 458 calcium alginate gel encapsulation, 495–6 calcium chloride hexahydrate, 72 calcium–cresolophthalein, 386 calculosis, 153 caldesmon, 340, 402, 403 calponin, 340, 402, 403 carbachol, 306, 308, 312, 324 carbon beads, 423 carbon buckyballs, 293 carbon nanotubes, 288, 293 carbonate-apatite, 158 catheter cross-sections of unused catheters, 183 crystalline biofilm formation, 163–4, 167 design, 179–83 and Proteus mirabilis biofilm formation, 157–85 encrustation main cause of complication, 162 prophylactic antimicrobials, 170–4 factors modulating rate of biofilm formation, 175–8 self-lubricating materials, 191–206 bioactive lubricious silicones, 201–2 biomimetic lubricious silicones, 203 self-lubricating silicone biomaterials, 197, 199 silicone chemistry, 192–7 toxicity and regulatory issues, 203–5 catheter-associated urinary tract infection, 26–7 cavernosography, 475, 476, 478 cavernosometry, 475, 476, 478 CD133, 520 cell fusion, 521

cell tracking technology, 349, 350–1 markers GFP, 350 LacZ, 350 others, 351 Y-chromosome, 350–1 methods, 349 ‘cell-based tissue-engineering,’ 526 cell’s gene library, 370 centronucleation, 448 C-Flex, 89 Chalk and Harrod method, 195 Change medium, 387 chicken ovalbumin, 72 chloroplatinic acid, 195 chlorosilanes, 192 ciprofloxacin, 138 claudin, 508, 524 cloning see reproductive cloning; therapeutic cloning CMMP-eGFP40, 384 coated ureteral stents carbon-coated stent analysis after 1 month of dwelling, 146–7, 150–1 elements ratio in brushite chemical formula, 150 elements ratio in calcium oxalate chemical formula, 150 FESEM of encrustation-free stent, 148 before insertion, 144 micrographs related to encrusted stent, 149 micro-IR analysis, 150–1 micro-IR spectra, 151 morphological characterisation before indwelling, 147 clinical study procedure, 137–9 discussion, 151–4 heparin-coated stent analysis, 139–41, 143–4 after 1 month indwelling, 139–41, 143–4 after prolonged indwelling, 144

Index electron microscopy characterisation before indwelling, 140 elements ratio in calcium oxalate chemical formula, 142 before insertion, 139 internal surface, 142 micro-IR spectra after 1 month indwelling, 143 micro-IR spectra after prolonged indwelling, 146 scanning electron micrographs after prolonged indwelling, 145 transversal cross-sections, 141 indications, 138 outcome assessment, 139–41, 143–4, 146–7, 150–1 collagen, 505, 540 as component of smart biomaterial, 248–9 extracellular matrices, 261–3, 265–74 acellular dermal matrix, 274–5 bladder extracellular matrix, 268–73 porcine small intestinal submucosa, 262–3, 265–7 as fundamental natural biomaterial, 259–61 for penile reconstruction, 473 collagenase type IV, 484 colony forming units, 20 colony-forming-unit fibroblasts, 397 Coloplast, 231, 233, 235 computed tomography axial, 123 micro, 314 ‘conditioning film,’ 7–8 Cook Urological, 30 Corning Glass Works of America, 193 corporal tissue, functional, engineering of, 473–6 Corvita endoluminal graft, 114 cryopreservation, 492–3 cTnI, 381 cTnT, 381

555

cystectomy, subtotal, 313 ‘cystocentesis,’ 49 cystoscopy, 438 cystourethroscopy, 435 cytokeratin family, 338 cytokines, as component of smart biomaterial, 250–1 cytotoxicity testing, 204–5 Dacron, 478 Danish Prostatic Symptom Score 1, 219 de-epidermised dermis, 457 ‘defensins,’ 23 Denis–Browne buried skin strip procedure, 464 desmin, 340, 387 detrusor-sphincter dyssynergia, 210–11 dextran, 289 dialkyltin dicarboxylates, 195 diamond-like coatings, 29 diethylenetriamine pentaacetic acid renography, 119 dog model as in vivo model for ureteral stent, 50–2 advantages and disadvantages, 52 kidney, 50 ureter and bladder, 50–1 urine biochemistry, 51 urolithiasis, 51–2 double pigtail stent, 111 metal, 119–21 double-J catheter, 90, 112, 115 polyurethane ureteral, 137 double-J stent, 91, 110, 111, 138 drug-eluting stents, 115–17 Duchene muscular dystrophy, 446 Dulbecco’s modified Eagle’s medium with 10% FBS, 484–5, 489, 490 with Ham’s F12, 488, 489 MCDB medium, 399 duloxetine, 423 elastin, 249 electrical field stimulation, 305, 307, 312, 313, 324

556

Index

electrocautery, 337 ‘electrophysiological signatures,’ 308 electrospinning, 286 Elipson analysis, 366 embryo cryopreservation, 492 embryoid bodies, 522 embryonic stem cells, 335, 361–3, 423, 521–2, 534–5 clinical applications, 362–3 ethical issues, 363 human, 533 overview, 361–2 properties, 362 for urological reconstruction, 357–71 encrustation BEST assay lid and assay plate, 76 of catheters main cause of complication, 162 prophylactic antimicrobials, 170–4 development of urinary encrustation, 62–4 dynamic continuous flow models, 73–6 dynamic flow-through models, 66–70 dynamic model representation, 75 dynamic recirculating encrustation model, 74 Finlayson and Dubois continuous flow vessels, 67 hydroxyapatite and struvite on Foley catheter, 64 MBEC-BEST assay, 76–7 models for assessment of formation on urological materials, 59–78 simple physical model of catheterised bladder, 69 static artificial urine model, 71 static models, 71–3 static vessel containing five stent sections, 72 temporary solution, 135–6 in vitro models, 65–6 artificial urine composition devised by Cox, 66 artificial urine composition devised by Griffith, 65

endoscopic therapy, 346–8 endothelial nitric oxide synthase, 410 end-stage renal disease, 503, 511 energy dispersive spectroscopy, 137 energy dispersive X-ray analysis, 71 enoyl-acyl carrier protein, 28 enoyl-acyl carrier protein reductase, 97 Enseal device, 293 eosin staining, 485 epigallocatechin-3-gallate, 27 ‘epigenetics,’ 370 Escherichia coli, uropathogenic, causing IBCs, 21–2 ethylene-vinyl-acetate, 89 external urethral sphincter, 425 extracellular matrix, 456 extracellular polymer substance, 9 7F double pigtail stent, 54 6F polyurethane stents, 91 Fabian stent, 211 spiral, 219 FBS see foetal bovine serum fibrin, 164 fibronectin, 505, 509 as component of smart biomaterial, 249 for penile reconstruction, 473 field emission scanning electron microscopy, 137 finasteride therapy, and biodegradable stents, 220 ‘fingerprint’ zone, 143, 144 Finlayson and Dubois model, 66–8 flow cytometry, 401 fluorofamide, 178–9 foetal bovine serum, 10% in K-SFM, 485 as supplement to DMEM, 484–5, 489, 490 Foley catheter, 64, 70, 136, 179–80, 182, 184, 214 folliculogenesis, 491 Fura-2, 307–8

Index galactocerebrosidase, 388 GATA4, 381 General Electric siloxane notation, 193 Geobacter sulfurreducens, 11 GFP, as biomarker for cell tracking, 350 Gleeson model, 67–9 glial-derived neurotrophic factor, 507–8, 524 glycocalyx-encased microcolonies, 16–17 glycogen synthase kinase-3 inhibitors, 323 glycosaminoglycans, as component of smart biomaterial, 250 Goretex, 478 G-protein-gated inwardly rectifying potassium, 388, 539 graciloplasty, 450 green fluorescent protein, 344–5, 384, 507 enhanced, 521 Grignard reaction, 193 growth factors, as component of smart biomaterial, 250–1 gyratory shaker, 77 gyrorotary platform, 73 Gyrus ACMI, 30 haematoxylin staining, 485 hAFSCs see human amniotic fluidderived stem cells heart failure, 446 hematopoietic stem cells, 380, 522–3 Hemobahn endoprosthesis, 113 hepatocyte growth factor receptor, 390 hepatocyte nuclear factor 4, 390 Higuchi equation, 202 histone deacetylase inhibitors, 323 HIV-Tat, 289 human amniotic fluid-derived stem cells, 505 human embryonic stem cells see embryonic stem cells hydrosilane, 195 hydroxyapatite, 63, 64, 72, 73, 75, 135 hypogastric nerves, 425

557

IBCs see intracellular biofilm communities iliotibial band, 458 immunomagnetic beads, 401 implants, tissue-engineered urological, performance assessment, 299–315 in vitro fertilisation, 492 in vitro maturation, 493 INCA X-Sight, 138 Incontinence Quality of Life, 436 ‘induced pluripotent state,’ 537–8 induced pluripotent stem cells, 369–70 inferior mesenteric ganglia, 425 inferior splanchnic nerves, 425 In-Flow, 181 infrared spectroscopy, 71, 137 inhibizone, 232 inkjet printing techniques, 473 insulin-like growth factor, 449 International Fat Applied Technology Society, 397 interstitial laser coagulation, 210, 219, 220 intracellular biofilm communities, caused by uropathogenic Escherichia coli, 21–2 Iscove’s modified Dulbecco’s medium, 507 with 2% foetal calf serum, 380 keratinocyte serum-free medium, 484, 485 kidney approaches for regeneration of tissue, 505–10 developmental approaches, 505–8 tissue engineering approaches, 508–10 basic components of tissue engineering, 503–5 biomaterials, 505 cells, 503–5 cell-based therapy for disease, 511–12

558

Index

regeneration, and stem cells, 518–27 regenerative medicine, 502–13 Klebsiella pneumoniae, 162 Krabbe globoid leukodystrophy, 388 K-SFM see keratinocyte serum-free medium LacZ as biomarker for cell tracking, 350 transfected hAFSCs, 507 Laminaria hyperborea, 257 Laminaria lessonia, 257 laminin, 505 leak point pressure, 432 leptin, 398 leukaemia inhibitory factor, 521–2 lipoblasts see adipose progenitor cells liposuction tumescent, 398 ultrasonic, 398–9 Loop Polaris, 95, 96 LPP see leak point pressure M199, supplemented with 10% foetal calf serum, 380 Madin–Darby canine kidney epithelial cells, 507 magnesium chloride hexahydrate, 73 magnetic resonance imaging, 117, 123, 124, 314 Matrigel, 386, 387, 389 Mayer–Rokitansky–Kuster–Hauser syndrome, 482 MBEC-BEST assay, 76–7 MDSC see muscle-derived stem cells Memokath, 211, 212 Memokath 028, 93, 211 Memokath 051, 114–15 mesenchymal amniocytes, ovine, 380–1 mesenchymal stem cell, 396, 523–4 from amniotic fluid, 379–82 classification criteria, 399 effect of mechanical stimulation, 403 vs adipose-derived stem cells, 401 metal stents applications, 109–18

balloon-expandable metal stents, 112–13 covered metal stents, 113–14 drug-eluting metal stents, 115–17 metal pigtail stent, 117–18 self-expandable metal stents, 109–112 thermo-expandable metal stents, 114–15 balloon dilatation of stenotic ureteral segment, 119 complications and problems, 121–3 encrustation, 122 migration, 122–3 urothelial hyperplasia, 121–2 excretory urography patent stented ureteral segment, 120 urothelial hyperplasia jeopardising luminal patency, 121 experience gained with various types of metal stents, 106–8 experimental studies in ureteral stenting, 127–8 extra-urinary drainage of upper urinary tract, 125–6 future trends, 126–7, 129 general types used in ureter, 105–6 insertion techniques, 118–21 conventional metal stent, 118–19 metal double-pigtail stent – Resonance, 119–21 properties of an ideal ureteral stent, 105 stented pig ureteral lumen fine epithelial lining covering metal mesh and lumen patency, 116 hyperplastic tissue expanding through the metal stent struts, 117 in the upper urinary tract, 104–29 and virtual endoscopy, 123–5 patent stented ureteral lumen, 124 stenotic distal part of endoprosthesis, 124

Index methyl cellulose, 182 methyl thiazolyel-tetrazolium, 510 metronidazole, 182, 202, 206 MHC see myosin heavy chains, skeletal microencapsulation, 493 micro-infrared spectrophotometer, 138 Microlance 3 needles, 72 minimum biofilm eradication concentration, 76 minimum essential medium elution test, 204–5 minocycline, 232 Mitomycin-C, 259 mixed cellular plus collagen injection therapy, 437 monolithic double-‘barreled’ stents, 30 Morganella morganii, 162 mouse embryonic fibroblast, 537 MPC see muscle precursor cell multidrug resistance membrane transporter, 390 multi-wall carbon nanotubes, 288 muscle precursor cell, 445 muscle-derived cell therapy, 428–9 muscle-derived stem cells, 423 myoblast transfer therapy, 428–9 myosin heavy chains, skeletal, 448 myosin light chain-2a, 381 myosin light chain-2v, 381 nanotechnology components of nano drug delivery system, 291 drug delivery system advantages over conventional methods, 290 components, 291 future trends, 292–3 manipulating nanomaterials, 293 materials safety, 293 smart nanomaterials, 292–3 nanomaterials for aiding cell tracking, 287–90 as biomaterials, 283–6 for drug delivery improvement, 290, 292

559

nanoscale system visualisation fluorescence microscopy, 287 gamma scintigraphy, 287 magnetic resonance imaging, 287 positron emission tomography, 287 single photon emission computed tomography, 287 ultrasound, 287 rationale for nanomaterials in engineering tissue, 282–3 in smart scaffolds development, 330 tissue scaffolds created using nanofabrication techniques, 285 and urological engineering, 281–93 ‘nanowires,’ 11 natural biomaterials for urological tissue engineering, 255–75 neobladder, 331 nestin, 381, 388 neural cell adhesion molecule, 449 neurofilament, 448 neuroleukin, 449 nitric oxide, as measure to evaluate detrusor contractility regulation, 307 norepinephrine, 426 notexin, 448 nuclear cloning, 363 nuclear transfer, 363–7 clinical applications, 364–7 ethical issues, 367 overview, 363–4 reproductive cloning, 364 types, 364 for urological reconstruction, 357–71 nuclear transplantation, 363 octamer-binding protein 4, 382, 383 1-octyl-2-dodecanol, 200 ‘off-the-shelf’ graft material, 455 scaffold, 540 Oil-O-Red, 385 oleyl alcohol, 182 oligonucleotides, 290

560

Index

Onuf’s nucleus, 426 Open-8 stent, 96 Open-I stent, 96 Open-Pass stent, 96 optical diffraction crystallography, 67 optical microscope, 137 organ bath system, 309 organosilicons, 192 organotin catalyst, 195 organotin silanoate, 197 osteocalcin, 386 osteopontin, 509 ovary methods for culturing follicles, 495–6 calcium alginate gel encapsulation, 495–6 culture media, 496 follicle isolation, 495 oocyte recovery and evaluation, 496 ovarian follicle encapsulated in calcium alginate bead, 494 tissue engineering, 491–5 cryopreservation, 492–3 folliculogenesis, 491–2 ovarian follicle tissue engineering, 493–5 Oxalobacter formigenes, 98 paclitaxel, 290 Palaquium gutta, 86 Palmaz-Schatz stent, 112, 113 ‘pancake’ reservoir, 231 parthenogenesis-derived stem cells, 367–9 clinical applications, 367–9 ethical issues, 369 overview, 367 for urological reconstruction, 357–71 Parylene coating, 234 Passager metal stent, 114 Paterson Forester stent, 8Fr, 125 pax-2, 521 pelvic nerves, 425 pelvic pain syndrome, chronic, 16 Pelvicol, 273

penile implants, 226–37 advances in design and materials used for inflatable implants, 228 biomaterials in current use, 230–1 cylinder materials, 230–1 design challenges of inflatable prosthesis, 230 reservoir materials, 231 device infection, 231–4 erosion resistance, 234–5 essential components fluid reservoir, 227 intra-penile paired cylinders, 227 valve pump, 227 future trends, 236–7 historical aspects of development, 227–30 hydrophilic coating, 233 inhibizone coating, 233 material characteristics of silicone and bioflex, 235 silicone sandwich with fabric and Parylene coating, 234 penile reconstruction, 470–9 basic principles of penile tissue engineering, 471–3 anatomy of penis, 472 delivery of cells, 473 scaffold design, 472–3 engineered penile prosthesis, 476–7 functional corporal tissue engineering, 473–6 summary and future trends, 478–9 tunica albuginea reconstruction, 478 penis anatomy, 472 implants (see penile implants) main cell types, 472 reconstruction (see penile reconstruction) Percuflex, 89 Percuflex Tail Plus, 95–6 pericytes see adipose progenitor cells ‘persisters,’ 6 Perspex tank, 71 PGA see polyglycolic acid

Index phenylephrine, in endogenous bladder regeneration, 326 phosphate-buffered saline, 485 pig model see swine model PKH26 green fluorescent cell linker, 387 platinum catalysed hydrosilation reaction, 192 platinum complex, 195 Pluorinc surfactant, 288 Polaris, 95 poly(anhydrides), 541 polydimethylsiloxane, 182, 192 polyglycolic acid, 456, 485, 541 polylactic acid, 541 poly(lactic-co-glycolic acid), 341, 407, 541 poly-L-lactic acid-coated polyglycolic acid, 477 polymerase chain reaction, 309, 365 polymyxin B, 182 poly(ortho-esters), 541 polypropylenimine dendrimers, 290 polysiloxane, 192 polytetrafluoroethylene, 423, 456 prazosin, in endogenous bladder regeneration, 324 preadipocytes see adipose progenitor cells processed lipoaspirate cells see adipose progenitor cells ProstaCoil, 211 Prostakath, 93, 181, 211, 218 prostatic obstruction, 210 prostatitis, chronic, 16 prosthesis, penile, engineered, 476–7 prosthetic valve endocarditis, 13–14 Proteus mirabilis biofilm formation and catheter design, 157–85 catheter design, 179–83 cross-sections of unused catheters, 183 crystalline biofilm formation on catheters, 163–4, 167

561

effect of triclosan-loaded stent, 98 electron micrographs of swimming normal cells, 160 epidemiology of infections, 162–3 factors modulating rate of biofilm formation on catheters, 175–8 biological control of encrustation, 176–8 nucleation pH of urine and rate of encrustation, 175–6 future trends, 183–4 petri dish showing swarming over catheter bridges and agar, 161 prophylactic antimicrobials, 170–4 mandelic acid, 171–2 nitrofurazone, 171 silver, 171 triclosan, 172–3, 174 scanning electron micrographs biofilm formation on hydrogelcoated latex catheters, 168 biofilms on catheters removed from laboratory models, 177 early stages of biofilm formation, 170 low-scanning, crystalline biofilms, 169 rafts of swarmer cells moving left to right, 161 surface of unused all-silicone catheters, 166 surface of unused latex-based catheters, 165 urease inhibitors, 178–9 virulence factors, 158–9, 162 migration over catheter surfaces, 159, 162 swarming, 158–9 urease, 158 Proteus vulgaris, 17 Providencia rettgeri, 162 Providencia vulgaris, 162 Pseudomonas aeruginosa, 9, 11, 162, 234 pudendal nerves, 425 pulsed-field gel electrophoresis, 163 pyelonephritis, 14–15

562

Index

quantum dots, 287–8 quantum phenomena, 283 QuickCore disposable biopsy needle, 435 quorom sensing compounds, 11, 28 rabbit model as in vivo model for ureteral stent, 47–50 advantages and disadvantages, 49– 50 calciuria and urolithiasis, 49 kidney, 47 ureteral microanatomy, 47–8 urinary bladder, 48 urine biochemistry, 48–9 radiograph diffraction crystallography, 67 Raman scattering, 288 Rapid-hyb buffer, 384 rat model as in vivo model for ureteral stent, 45, 47 advantages and disadvantages, 47 bladder, 45, 47 kidney, 45 ureteral microanatomy, 45 urine biochemistry, 45 Recombinant human bFGF, 387 ‘reconstituted collagen,’ 249 ‘red rubber catheter,’ 87 regenerative medicine, 322–3 see also tissue engineering of kidney, 502–13 urinary sphincter via an endoscopic approach, 422–38 of urinary sphincter via direct injection, 445–51 urological reconstruction conclusion and future trends, 370–1 embryonic stem cells, 361–3 induced pluripotent stem cells, 369–70 nuclear transfer, 363–7

parthenogenesis, 367–9 stem cells, 360–1 regenerative pharmacology aim, 323 application, 324 and bladder regeneration, 322–32 critical role in regenerative medicine and tissue engineering, 325 definition, 323 renal cell therapy, 511 reproductive cloning, 364, 535–6 reproductive medicine tissue engineering, 482–97 reprogramming, 537–8 direct, 369 ‘resist,’ 233 resistin, 398 Resonance, 90, 91 double-pigtail metal stent, 115, 117–18 resveratrol, 179 retention balloon, 171 retrograde urethral perfusion pressure, 409–10 reverse transcriptase polymerase chain reaction, 507, 509, 510 rhabdosphincter, 425 rifampin, 232 Robbins device, modified, 73 room temperature vulcanisation, 193 Russell’ technique, 464 sarcometic tropomysin, 387 scaffold technology cell-seeded, 341 vs non-seeded, 341–2 Schewanella oneidensis, 11 scopolamine, in endogenous bladder regeneration, 324 self-expandable permanent endoluminal stent, 90 serotonin, 426 Serratia marcescens, 4 silanolysis, 197 silastic catheter, 509 silica skeletons, 164, 167

Index silicone, 192, 456 alkoxy crosslinking mechanism, 196 bioactive lubricious, 201–2 as biocompatible stent material, 89 biomimetic lubricious, 203 chemistry, 192–7 commonly substituted groups with their applications, 194 crosslinking for elastomer production, 193–5, 197 ethylene bridge formation, 195 General Electric siloxane notation, 194 history, 192–3 nomenclature, 193 platinum catalysed hydrosilation mechanism, 196 conventional and novel crosslinkers, 197 effect of benzalkonium chloride patency solution, 204 lubricious and standard RTV silicone elastomer static and dynamic coefficients of friction, 200 tensile properties, 201 particles, 423 performance characteristics, 199–201 apparatus for friction coefficient determination, 199 determination of coefficients of friction, 199–200 mechanical strength determination, 200–1 polydimethylsiloxane chemical structure, 194 self-lubricating, 182 self-lubricating materials, 191–206, 197, 199 tetra-n-alkoxysilane crosslinkers, 198 toxicity and regulatory issues, 203–5 cytotoxicity testing, 204–5 elastomer biocompatibility, 203–4 Silitek, 89 single-wall carbon nanotubes, 288 SIS see small intestinal submucosa

563

‘slime,’ 63 SM-22, 402, 403 SM myosin heavy chain, 402, 403 small intestinal submucosa, 345, 459–60, 542 composition adhesive glycoproteins, 459 fibrillar collagen, 459 layers lamina propria, 459 muscularis mucosa, 459 tunica submucosa, 459 porcine, 262–3, 265–7 comparison of physical structures, 263, 265 components, 263 fundamentals, 262–3, 265 poor regeneration, 266 pre-clinical studies, 265–7 preparation technique, 262 uses, 265, 267 smart nanomaterials, 292–3 ‘smart scaffolds,’ 248 smooth muscle induction media, 403 smooth muscle myosin heavy chain, 340 smoothelin, 339–40, 402, 403 Sof-Curl, 95 somatic cell nuclear transfer see also therapeutic cloning and tissue engineering, 525–6 Song urethral stent, 212 Southern blot analysis, 384 Spanner, 214 sphincterotomy, transurethral, 210 spiral computed tomography, 123 Spirastent, 92 SpiroFlow, 93, 181 stage-specific embryonic antigen-3, 383 stage-specific embryonic antigen-4, 383 staghorn renal calculi, 17 Staphylococcus aureus, 234 Staphylococcus epidermidis, 232, 234 static models, 71–3

564

Index

Statistica, 138 stem cell therapy cell types adult stem cells, 423 embryonic stem cells, 423 vs collagen injections, 436–7 stem cells, 360–1 endogenous, 519–21 bone marrow-derived stem cells, 520 cell fusion, 521 tissue-specific progenitor/stem cells, 520–1 exogenous, 521–6 amniotic fluid/placental stem cells, 524 bone marrow-derived stem cells, 522–4 embryonic stem cells, 521–2 somatic cell nuclear transfer and tissue engineering, 525–6 and kidney regeneration, 518–27 multipotent, 360 pluripotent, 360 sources, 361 totipotent, 360 unipotent, 360 stents see also specific stents arising problems, 88 coatings, 93–5 glycosaminoglycan, 94 heparin, 94–5 hydrogel, 93–4 monomethoxy poly(ethyleneglycol)-3,4dihydroxyphenylalanine, 94 oxalate decarboxylase, 94 pentosan, 94 phosphorylcholine, 94 plasma-deposited diamond-like coatings, 94 polyvinyl pyrrolidone, 94 current biomaterials, 88–93 bare metal stents, 91–2 biodegradable stents, 92–3 metal ureteral stents, 90–1 new materials, 89

definition, 86 design, 95–6 bladder curl, 95 renal curl, 95 drug-eluting stents, 97–8 flow areas of 7F stents, 97 future trends, 98–9 history of biomaterials, 86–8 metal (see metal stents) metallic, 211–14 clinical outcome, 212–13 covered, 212 first-generation, 211 second-generation, 211–12 temporary, associated complications, 213–14 Open-pass stent, 96 Resonance metal stent, 90 ureteral coated, 134–54 design and materials, 85–99 ′I-beam configuration of Openand Open-I, 97 pigtail curls, 86 urethral (see urethral stent) uses in urology, 85 Strecker stent, 112, 113 stress urinary incontinence, 345, 445 adult stem cell therapy efficacy, 437 autologous stem cell injection therapy, 424 categories, 423 risk factors, 422 current results of clinical studies, 436–8 injection technique, 434–6 endoscopic injection technique, 435–6 percutaneous needle muscle biopsy, 434–5 MDSCs in delivery of neurotrophic factors, 433–4 neurophysiology, 424–6 bladder and urethral responses during sneezing, 427

Index peripheral nerves controlling urethral muscles, 425 urethral pressure response, 428 urogenital tract innervation, 425 role of pharmacotherapy, 423 stem cell source for injection therapy, 427–9, 431–3 adipose-derived stem cells, 428, 433 bone marrow stroma, 427 cauterised mid-urethra after Hank’s balanced salt solution, 431 human MDSC proliferating in growth culture medium, 429 LPP difference after MDSC injection, 433 muscle-derived stem cells, 428–9, 431–3 myotubes and myofibers from MDSCs, 430 striated muscle layer and innervation of cauterised midurethra, 432 stricturotomy, 464 struvite, 63, 64, 72, 73, 75, 135, 158, 175, 182, 270 suburethal sling, 438 superparamagnetic iron oxide, 288, 289 surface antigen c-Kit, 383 surfactant benzalkonium chloride, 203 ‘Sushruta Samhita,’ 87 swarming, 158–9 swine model as in vivo model for ureteral stent, 52–4 advantages and disadvantages, 53–4 arterial and venous arrangements, 53 kidneys, 52 ureter microanatomy, 52–3 urine biochemistry, 53 symphathetic chain ganglia, 425

565

Tamm-Horsfall protein, 23, 511 TA-XT2 Texture Analyser, 201 TAXUS, 116 Tecoflex, 89 teicoplanin, 28 tensor fascia lata, 458 teratocarcinomas, 383 tetraethylsilane, 192 tetraoctyldodecoxysilane, non-volatile, 197 tetraoleyloxysilane, 199 tetrapropoxysilane, 197 therapeutic cloning, 364, 535–7 thermal inkjet printing, 386 Thermus aquaticus, 4 thiazolidinediones, 397 tissue engineering biomaterials, 539–41 acellular tissue matrices, 541 naturally derived materials, 540 synthetic polymers, 541 bladder construction of tissue or organ, 327 endogenous bladder regeneration, 324, 326–7 implantation in preclinical studies, 330–1 preliminary clinical experience with neobladders, 331 and regenerative pharmacology, 322–32 bladder repair and replacement, 541–3 bladder replacement, 543 matrices for bladder regeneration, 542–3 seromuscular grafts and de-epithelialised bowel segments, 542 tissue expansion for bladder augmentation, 542 cell-based allogeneic, 335 autologous, 335 xenogeneic, 335

566

Index

cells used, 534–9 embryonic stem cells, 534–5 native cells, 534 placental and amniotic fluid stem cells, 538–9 reprogrammed somatic cells, 537–8 therapeutic cloning, 535–7 comparison of trigone-sparing cystectomy results, 544 construction of engineered human bladder, 544 penile, basic principles, 471–3 principles, 358–60 in reproductive medicine, 482–97 method for culturing follicles, 495–6 ovarian tissue, 491–5 uterine tissue reconstitution, 487– 91 uterus, 486–7 vagina, 482–3 vaginal tissue reconstitution, 483–5 results of tissue engineered bladder implantation, 545 and somatic cell nuclear transfer, 525–6 strategies use of acellular matrices, 533 use of matrices seeded with cells, 533 techniques classification, 359 techniques for bladder tissue, 532–45 and therapeutic cloning strategies, 536 urological implants performance assessment, 299–315 urological reconstruction conclusion and future trends, 370–1 embryonic stem cells, 361–3 induced pluripotent stem cells, 369–70 nuclear transfer, 363–7 parthenogenesis, 367–9 stem cells, 360–1

tissue-engineered buccal mucosa, 457 Titan implants, 233 Tra-1-81, 383 transexualism, 482 transferrinlipoplex, 290, 292 transurethral microwave thermotherapy, high-energy, 210, 219 transvaginal tape, 438 Trestle catheter, 214 triclosan, 27–8, 97, 162, 172–3, 174 tritiated thymidine, 510 Triumph, 28 tropocollagen, 248 trypsin-ethylenediaminetetraacetic acid, 485 tryptic soy broth, 68 TUDS, 93 tunica albuginea, reconstruction, 478 type 1 fimbriae FimH receptors, 23 urease, 158 inhibitors, 178–9 Ureteral Stent Symptom Questionnaire, 88 ureteral stents coated, 134–54 commonly used animal models, 45– 54 in vivo models, 42–55 history, 42–3 vs in vitro models, 43–4 ureteropyelography, retrograde, 138, 139 urethra biological (natural) scaffolds, 456–64 acellular dermal matrix, 457 amniotic membrane, 458–9 bladder acellular matrix graft, 460–4 cadaveric fascia lata, 458 small intestine submucosa, 459–60 regenerative medicine, 454–65 synthetic scaffolds, 455–6 tissue engineering using a collagen matrix, 462

Index urethral catheter gentamicin-releasing, 136 history, 60 urethral stent biodegradable, 214–21 bacterial adherence, 218 bioabsorption and biodegradation, 215 biocompatibility, 217 biodegradable materials, 215–16 braided PLA, 217 degradation times of different bioabsorbable copolymers, 218 encrustation, 218 mechanical properties and degradation of self-reinforced composite devices, 216–17 prostatic stents, 218–20 requirements, 214–15 for urethral strictures, 220–1 future trends, 221–2 indications for use, 209–11 detrusor-sphincter dyssynergia, 210–11 prostatic obstruction, 210 urethral stricture, 209–10 nondegradable temporary, 211–14 metallic stents, 211–14 polymer stents, 214 temporary, 208–22 urethral stricture main types, 209 use of temporary urethral stents, 209–10 uricase enzyme, 52 urinary sphincter, regenerative medicine via an endoscopic approach, 422–38 via direct injection, 445–51 challenges with MPC transfer, 446–7 conclusions and future trends, 451 direct myofibre implantation procedure, 447–50

567

MHC and neurofilament immunostaining, 449 of MPC using minced mice, 450 urinary stress incontinence autologous progenitor cell-based therapy, 347–8 urine-derived progenitor cells, 342–4 cost, 344 potential advantages, 343 UroCoil, 212 urological engineering and nanotechnology, 281–93 urological reconstruction autologous cell sources, 334–51 cell tracking technology, 348, 350–1 embryonic stem cells, nuclear transfer and parthenogenesisderived stem cells, 357–71 fully differentiated cells, 337–42 from non-urological origin, 340–1 from urological tissues, 337–40 uses in urinary tract reconstruction, 341–2 stem/progenitor cells, 342–8 autologous, for endoscopic therapies, 346–8 methods of inducing differentiation, 345–6 from non-urological tissues, 344–5 from urinary tract, 342–4 urological tissue engineering artificial biomaterials, 243–52 implants, performance assessment, 299–315 natural biomaterials, 255–75 uroplakins, 338 Urospiral, 181, 211 uterus endometrial cells isolated from uterine tissue, 490 smooth muscle cells isolated from uterine tissue, 490 tissue engineering, 486–7 tissue reconstitution methods, 487–91 cell characterisation, 489–90

568

Index

cell seeding and construct formation, 490 culture medium, 489 histology, molecular, tensile strength and organ bath studies, 491 protocols for culturing smooth muscle cells, 489 protocols for isolation and culturing of endometrial cells, 488–9 unseeded and seeded PGA/PLGA construct, 491 vagina keratinocytes growing out of a piece of vaginal tissue, 484 methods of tissue reconstitution, 483–5 cell characterisation, 485 cell isolation and culture, 483–4 culture medium, 484–5 histology, molecular, tensile strength and organ bath studies, 485 scaffold construction and cell seeding, 485 subculture of the cells, 485 PGA scaffold seeded with vaginal epithelial and smooth muscle cells, 486 tissue engineering, 482–3 vesicoureteral reflux cell-based therapy, 347

myocyte- or stem cell-based injection, 347 vesicular acetylcholine transporter, 448 Vibrio fischeri, 9 Vicryl, 216 vimentin, 340, 521 virtual endoscopy, and metal stents, 123–5 limitations, 124–5 main goal, 123 visual laser ablation of the prostate, 210, 219 von Kossa staining, 386 vulcanisation, of silicones for elastomer production, 193–5, 197 main reaction mechanisms condensation or addition cures, 194–5 hydrosilation crosslinking, 195 Wallstent, 90, 109, 110, 113, 114, 122 self-expandable metal stent, 54 Western blot analysis, 309, 345, 485 Wolffian duct, 505 wound infections, 13 X-ray diffraction spectroscopy, 71 Y-chromosome, as biomarker for cell tracking, 350–1 zona occluden-1, 507, 524