Tissue Engineering of Vascular Prosthetic Grafts

TISSUE ENGINEERING INTELLIGENCE UNIT 1

Tissue Engineering of Vascular Prosthetic Grafts Peter Zilla, M.D., Ph.D. University of Cape Town Cape Town, South Africa

Howard P. Greisler, M.D. Loyola University Maywood, Illinois, U.S.A.

R.G. LANDES COMPANY AUSTIN, TEXAS U.S.A.

TISSUE ENGINEERING INTELLIGENCE UNIT Tissue Engineering of Vascular Prosthetic Grafts R.G. LANDES COMPANY Austin, Texas, U.S.A.

Copyright © 1999 R.G. Landes Company All rights reserved. No part of this book may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopy, recording, or any information storage and retrieval system, without permission in writing from the publisher. Printed in the U.S.A. Please address all inquiries to the Publishers: R.G. Landes Company, 810 South Church Street, Georgetown, Texas, U.S.A. 78626 Phone: 512/ 863 7762; FAX: 512/ 863 0081 ISBN: 1-57059-549-6

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Library of Congress Cataloging-in-Publication Data

Tissue engineering of prosthetic vasxcular grafts [edited by] Peter Zilla, Howard P. Greisler. p. cm.--(Tissue engineering intelligence unit) Includes bibliographical referecnces and index. ISBN 1-57059-549-6 (alk. paper) 1. Blood vessel prosthesis. 2. Vascular grafts. 3. Biomedical materials. I. Zilla, P.P. (Peter Paul) II. Greisler, Howard P. III. Series. [DNLM: 1. Blood Vessel Prosthesis. 2. Blood Vessels--transplantation. 3. Cell Transplantation. 4. Endothelium, Vascular--cytology. 5. Host vs Graft reaction--immunology. 6. Biocompatyible Materials. WG 170 T616 1999] RD598.55.T57 1999 617.4’130592--dc21 DNLM/DLC 99-24890 for Library of Congress CIP While the authors, editors and publisher believe that drug selection and dosage and the specifications and usage of equipment and devices, as set forth in this book, are in accord with current recommendations and practice at the time of publication, they make no warranty, expressed or implied, with respect to material described in this book. In view of the ongoing research, equipment development, changes in governmental regulations and the rapid accumulation of information relating to the biomedical sciences, the reader is urged to carefully review and evaluate the information provided herein.

CONTENTS PART I Bio-Inert Prostheses: Insufficient Healing 1. The Lack of Healing in Conventional Vascular Grafts ........................................................................................... 3 Lester Davids, Terri Dower, Peter Zilla

Introduction ..................................................................................................................................................................... 3 Midgraft Healing .............................................................................................................................................................. 4 ePTFE Grafts ..................................................................................................................................................................... 5 Dacron Grafts ................................................................................................................................................................... 9 Transanastomotic Healing ............................................................................................................................................. 15 Pannus Tissue ................................................................................................................................................................. 15 Transanastomotic Endothelialization ........................................................................................................................... 17 Porosity ........................................................................................................................................................................... 18 Fibrin ............................................................................................................................................................................... 26 Macrophages ................................................................................................................................................................... 30 Prosthetic Wall Vascularization ..................................................................................................................................... 34 Conclusion ...................................................................................................................................................................... 37

2. Noncompliance: The Silent Acceptance of a Villain ............................................................................................. 45 Alexander M. Seifalian, Alberto Giudiceandrea, Thomas Schmitz-Rixen, George Hamilton

Introduction ................................................................................................................................................................... 45 Historical Overview ........................................................................................................................................................ 45 Physical Properties of the Vessel Wall ........................................................................................................................... 46 Assessment of Compliance ............................................................................................................................................ 46 Synthetic and Biological Grafts ..................................................................................................................................... 49 Causes of Graft Failure ................................................................................................................................................... 50 Tubular Compliance ....................................................................................................................................................... 51 Anastomotic Compliance Mismatch ............................................................................................................................. 51 Signal Transduction Pathways and Flow ....................................................................................................................... 52 Development of a Graft with Better Compliance ......................................................................................................... 52 The Future ...................................................................................................................................................................... 55 PART II

Biolized Prostheses: Surface Healing 3. Endothelial Cell Seeding: A Review ...................................................................................................................... 61 Steven P. Schmidt, Gary L. Bowlin

Introduction ................................................................................................................................................................... 61

4. Surface Precoating in the 1980s: A First Taste of Cell-Matrix Interactions .......................................................................................................... 69 J. Vincent Smyth, Michael G. Walker

Endothelial Cell Attachment to Prosthetic Grafts ........................................................................................................ 70 Cell Migration ................................................................................................................................................................ 72 Endothelial Cell Retention Under Flow Conditions .................................................................................................... 72 Seeding of Native Vascular Surfaces .............................................................................................................................. 73 Mechanism of Cell Adherence ....................................................................................................................................... 73 Nonendothelial Cell Lines .............................................................................................................................................. 75 Future Directions ........................................................................................................................................................... 75

5. Surface Precoating in the 1990s: The Fine Tuning of Endothelial Cell Transplantation ..................................................................................... 79 Mark M. Samet, Victor V. Nikolaychik, Peter I. Lelkes

Introduction ................................................................................................................................................................... 79 Endothelialization: Past and Current Approaches ....................................................................................................... 80 Fine Tuning of Endothelialization via Surface Coating ............................................................................................... 80 Morphological Aspects of the Endothelialized Surfaces .............................................................................................. 84 Concluding Remarks ...................................................................................................................................................... 86

Microvascular Endothelial Cell Transplantation ..................................................................................91 6. Microvascular Endothelial Cell Transplantation: A Review ................................................................................ 93 Stuart K. Williams

Early Development of Endothelial Cell Transplantation Technology ........................................................................ 93 Cultured Endothelium for Transplantation ................................................................................................................. 94 Current Considerations in Endothelial Cell Transplantation ..................................................................................... 94 Markers for Isolated Endothelial Cells .......................................................................................................................... 95 Cell Deposition on Graft Surfaces: Seeding vs. Sodding .............................................................................................. 96 Animal Models of Endothelial Cell Transplantation. .................................................................................................. 97 Mechanisms Underlying the Formation of a Neointima in Tissue Engineered Vascular Grafts Using Microvascular Endothelial Cell Transplantation ............................................................................................ 97 Future of Microvessel Endothelial Cell Transplantation ............................................................................................. 99

7. Morphological Aspects of Microvascular Cell Isolates ...................................................................................... 101 Manuela Vici

Introduction ................................................................................................................................................................. 101 Microvascular Cells ...................................................................................................................................................... 102 Electron Microscopic and Immunocytochemical Profiles of Microvascular Cell Isolates ...................................... 103 Discussion ..................................................................................................................................................................... 110

8. Functional Aspects of Microvascular Cell Isolates ............................................................................................. 115 M. Fittkau, Teddy Fischlein

Antiaggregatory Properties of Microvascular Endothelial Cells ............................................................................... 116 Anticoagulative Properties ........................................................................................................................................... 116 Profibrinolytic and Antifibrinolytic Properties ......................................................................................................... 117 Conclusion .................................................................................................................................................................... 119

9. Automated Seeding Devices ................................................................................................................................ 121 Dominic Dodd, J. Vincent Smyth, Michael G. Walker

Introduction ................................................................................................................................................................. 121 Early Seeding Techniques ............................................................................................................................................. 122 Optimal Conditions for Seeding ................................................................................................................................. 122 Choice of Seeding Technique ....................................................................................................................................... 123 Summary ....................................................................................................................................................................... 125

10. Healing Patterns Following Microvascular Seeding— A Clinical Evaluation of Microvascular-Seeded A-V Access Grafts ............................................................... 127 Steven P. Schmidt, Sharon O. Meerbaum, Duane L. Donovan

Introduction ................................................................................................................................................................. 127 Materials and Methods ................................................................................................................................................ 128 Results ........................................................................................................................................................................... 129 Discussion ..................................................................................................................................................................... 129

11. Neointimal Hyperplasia in Small Diameter Prosthetic Vascular Grafts: Influence of Endothelial Cell Seeding with Microvascular Omental Cells in a Canine Model .................................................................................................................... 131 Miralem Pasic, Werner Müller-Glauser, Marko Turina

Introduction ................................................................................................................................................................. 131 Experimental Studies with Endothelial Cell Seeding at the University Hospital Zurich ......................................... 132 Endothelial Cells Reduce Late Neointimal Proliferation ........................................................................................... 134 Intimal and Neointimal Hyperplasia .......................................................................................................................... 137 Omental Microvascular Cells ...................................................................................................................................... 139 Limitations of Experimental Studies ........................................................................................................................... 140

12. Human Clinical Trials of Microvascular Endothelial Cell Sodding .................................................................. 143 Stuart K. Williams

History of EC Transplantation .................................................................................................................................... 143 Human Trial of Microvascular Endothelial Cell Transplantation ............................................................................. 144 Conclusions .................................................................................................................................................................. 147

Macrovascular Endothelial Cell Transplantation ................................................................................149 13. In Vitro Endothelialization: Its Contribution Towards an Ideal Vascular Replacement ............................................................................. 151 Peter Zilla

Historical Perspective ................................................................................................................................................... 152 Current and Future Perspective ................................................................................................................................... 153

14. Serial Cultivation of Human Endothelial Cells .................................................................................................. 157 Caroline Gillis-Hægerstrand, Anders Hægerstrand

Introduction ................................................................................................................................................................. 157 Characteristics of the Endothelium ............................................................................................................................ 157 Sources of Endothelial Cells ........................................................................................................................................ 158 Culture Techniques for Human ECs ........................................................................................................................... 158 Does the Culture Technique Matter from a Clinical Perspective? ............................................................................. 159 Future Possibilities of Cultured ECs ........................................................................................................................... 160

15. Risk Factors for Autologous Endothelial Cell Cultures ..................................................................................... 163 Johann Meinhart, Manfred Deutsch, Peter Zilla

Introduction ................................................................................................................................................................. 163 Standard Cell Culture Technique for Autologous Endothelial Cell Cultures ........................................................... 163 Procedure Related Risk Factors ................................................................................................................................... 164 Patient Related Risk Factors ........................................................................................................................................ 165 Summary and Future Aspects ...................................................................................................................................... 166

16. Adhesion Molecule Expression Following In Vitro Lining ................................................................................ 171 Caroline Gillis-Hægerstrand

Introduction ................................................................................................................................................................. 171 Vascular Grafts .............................................................................................................................................................. 171 Autologous Grafts ......................................................................................................................................................... 171 Prosthetic Grafts ........................................................................................................................................................... 171 Glutaraldehyde-Tanned Grafts .................................................................................................................................... 172 Heart Valve Prostheses ................................................................................................................................................. 172 Inflammation ................................................................................................................................................................ 172 Leukocyte Adhesion Under Flow Conditions ............................................................................................................. 173 Endothelialization Reduces Monocyte Adhesion to Xenogeneic Tissue in a Time Dependent Manner ............................................................................................... 174 Future Prospect: Antiadhesive Therapy? .................................................................................................................... 175 Concluding Remarks .................................................................................................................................................... 175

17. In Vitro Endothelialization Elicits Tissue Remodeling Emulating Native Artery Structures ................................................................................................................ 179 Manfred Deutsch, Johann Meinhart, Peter Zilla

Case Report ................................................................................................................................................................... 180 Conclusions .................................................................................................................................................................. 183

18. In Vitro Endothelialization of Synthetic Vascular Grafts in Long Term Clinical Use ................................................................................................................................ 189 Manfred Deutsch, Johann Meinhart, Peter Zilla

Introduction ................................................................................................................................................................. 189 Laboratory Procedure .................................................................................................................................................. 189 Surgical Procedure and Clinical Follow Up ................................................................................................................ 190 Randomized Clinical Study ......................................................................................................................................... 190 Clinical Routine Endothelialization ............................................................................................................................ 191 Conclusion .................................................................................................................................................................... 192

PART III Biointeractive Prostheses: Complete Healing Biological Components Taming of Adverse Responses ...................................................................................................................... 197 Prevention of the Inflammatory Reaction ............................................................................................................ 197 19. Inflammatory Reaction: The Nemesis of Implants ............................................................................................ 197 James M. Anderson

Introduction ................................................................................................................................................................. 197 Inflammation and the Healing Response .................................................................................................................... 197 Foreign Body Reaction ................................................................................................................................................. 199 Macrophage Motility, Adhesion and Activation ......................................................................................................... 201 Foreign Body Giant Cell Formation ............................................................................................................................ 204 Future Perspectives on Inflammatory Responses to Tissue Engineered Prosthetic Vascular Grafts ..................................................................................................... 205

Taming of Adverse Responses ...................................................................................................................... 207 Prevention of the Inflammatory Reaction ............................................................................................................ 207 20. Molecular Determinants of Acute Inflammatory Responses to Biomaterials ................................................. 207 Liping Tang, John W. Eaton

Introduction ................................................................................................................................................................. 207 Surface-Protein Interactions ........................................................................................................................................ 207 Fibrin(ogen) Is Necessary for Phagocyte Accumulation on Biomaterial Implants .................................................. 209 The Mechanism of Biomaterial-Mediated Inflammatory Responses ....................................................................... 212 Conclusions .................................................................................................................................................................. 214

Taming of Adverse Responses ...................................................................................................................... 219 Prevention of Fibrosis ............................................................................................................................................ 219 21. The Accumulation of Inflammatory Mediators: A Target for the Prevention of Fibrosis ........................................................................................................... 219 John Zagorski, Sharon M. Wahl

Introduction ................................................................................................................................................................. 219 Inflammation ................................................................................................................................................................ 220 Regulation of Leukocyte Recruitment ........................................................................................................................ 220 Mitigation of Recruitment ........................................................................................................................................... 223 Summary and Perspectives .......................................................................................................................................... 224

Taming of Adverse Responses ...................................................................................................................... 229 Prevention of a Hyperplastic Intimal Response ................................................................................................... 229 22. Pathobiology of Hyperplastic Intimal Responses .............................................................................................. 229 Erik L. Owens, Alexander W. Clowes

Introduction ................................................................................................................................................................. 229 Arteries .......................................................................................................................................................................... 230 Grafts ............................................................................................................................................................................. 235 Summary ....................................................................................................................................................................... 236

Taming of Adverse Responses ...................................................................................................................... 241 Prevention of a Hyperplastic Intimal Response ................................................................................................... 241 23. Cell Cycle Interruption to Inhibit Intimal Hyperplasia .................................................................................... 241 Michael J. Mann, Ruediger C. Braun-Dullaeus, Victor J. Dzau

The Molecular and Cellular Biology of Neointimal Vascular Graft Disease............................................................. 242 Cell Cycle Progression: A Careful Orchestration ....................................................................................................... 243 Cell Cycle Arrest and Neointimal hyperplasia ............................................................................................................ 245 Genetic Engineering of Vein Graft Resistance to Atherosclerosis via Cell Cycle Gene Blockade .................................................................................................................................... 246 Summary ....................................................................................................................................................................... 247

Facilitation of Healing.................................................................................................................................. 251 Chemotaxis ............................................................................................................................................................ 251 24. Signaling Mechanisms for Vascular Cell Migration ........................................................................................... 251 Ian Zachary

Introduction ................................................................................................................................................................. 251 Migration Factors ......................................................................................................................................................... 252 Future Perspectives ....................................................................................................................................................... 262

Facilitation of Healing.................................................................................................................................. 271 Chemotaxis ............................................................................................................................................................ 271 25. Adhesion Molecules: Potent Inducers of Endothelial Cell Chemotaxis ........................................................... 271 Zoltan Szekanecz, Alisa E. Koch

Introduction ................................................................................................................................................................. 271 The Role of Extracellular Matrix Macromolecule in Endothelial Cell Motility During Vessel Formation ........................................................................................... 272 Adhesion Molecules in Endothelial Cell Migration and Angiogenesis ..................................................................... 272 Interactions of Soluble Mediators with Cellular Adhesion Molecules and Extracellular Matrix Components During Endothelial Cell Recruitment and Neovascularization ............................................................................................................................................. 273 The Relevance of Angiogenesis Studies in Vascular Surgery: Aortic Aneurysms and Wound Healing .................................................................................................................... 274 A Regulatory Network in Sites of Endothelial Cell Migration: Potential Target Promoting Graft Healing ............................................................................................................... 275

Facilitation of Healing.................................................................................................................................. 279 Angiogenesis ........................................................................................................................................................... 279 26. Angiogenesis in Tissues and Vascular Grafts ...................................................................................................... 279 Paula K. Shireman, Howard P. Greisler

Overview of Angiogenesis ............................................................................................................................................ 279 Angiogenesis in Vascular Grafts .................................................................................................................................. 282 Therapeutic Angiogenesis in Ischemic Tissues ........................................................................................................... 283 Future Directions ......................................................................................................................................................... 284

Facilitation of Healing.................................................................................................................................. 287 Angiogenesis ........................................................................................................................................................... 287 27. Polypeptide Growth Factors with a Collagen Binding Domain: Their Potential for Tissue Repair and Organ Regeneration........................................................................... 287 Bo Han, Lynn L.H. Huang, David Cheung, Fabiola Cordoba, Marcel Nimni

Introduction ................................................................................................................................................................. 287 The Collagen Derived Matrix ...................................................................................................................................... 287 Interactions of Biosynthetic Matrices with Cells ....................................................................................................... 289 Modified Growth Factors: TGF-β With a Collagen Binding Domain ....................................................................... 291 Summary and Conclusion ........................................................................................................................................... 297

Facilitation of Healing.................................................................................................................................. 301 Angiogenesis ........................................................................................................................................................... 301 28. Fibroblast Growth Factors in Angiogenesis and Tissue Engineering ............................................................... 301 Karin A. Blumofe, Timothy J. Heilizer, Paula K. Shireman, Howard P. Greisler

Introduction ................................................................................................................................................................. 301 History .......................................................................................................................................................................... 301 Angiogenesis ................................................................................................................................................................. 302 FGF Characterization ................................................................................................................................................... 303 Structure ....................................................................................................................................................................... 303 FGF Secretion ............................................................................................................................................................... 303 Interaction with Heparin ............................................................................................................................................. 304 FGF and Vascular Grafts .............................................................................................................................................. 307 FGF and Ischemia ......................................................................................................................................................... 309 Conclusions .................................................................................................................................................................. 309

Facilitation of Healing .................................................................................................................................. 313 Matrix Degradation .............................................................................................................................................. 313 29. Role of Urokinase-Type Plasminogen Activator (uPA) in In Vitro Angiogenesis in Fibrin Matrices .................................................................................................... 313 Pieter Koolwijk, Victor W.M. van Hinsbergh

Summary ....................................................................................................................................................................... 313 Introduction. ................................................................................................................................................................ 313 Components of the Plasmin/Plasminogen Activator System .................................................................................... 314 The Regulation of Plasminogen Activation and Pericellular Proteolysis by Inflammatory and Angiogenic Mediators ........................................................................................................... 316 Interaction Between the uPA/Plasmin System and Matrix-Degrading Metalloproteases ....................................... 316 Involvement of uPA and uPA Receptor in Cell Migration and Realignment of Endothelial Cells ...................................................................................................................... 317 Role of the Plasmin/Plasminogen Activator System in the Formation of Endothelial Tubes in a Fibrin Matrix ...................................................................................... 318 Perspective .................................................................................................................................................................... 320

Facilitation of Healing .................................................................................................................................. 325 Matrix Modulation ................................................................................................................................................ 325 30. How Does Extracellular Matrix Control Capillary Morphogenesis? ................................................................ 325 Robert B. Vernon, E. Helene Sage

Angiogenesis: An Introduction .................................................................................................................................... 325 Models of Angiogenesis: The Essential Role of ECM ................................................................................................. 326 Interactions Between ECs and ECM that Mediate Angiogenesis .............................................................................. 330

Facilitation of Healing .................................................................................................................................. 343 Matrix Modulation ................................................................................................................................................ 343 β .......................................................................... 343 31. Collagen Matrices Attenuate Fibroblast Response to TGF-β Richard R. Clark, John M. McPherson

Introduction ................................................................................................................................................................. 343 Results ........................................................................................................................................................................... 344 Discussion ..................................................................................................................................................................... 348

Facilitation of Healing .................................................................................................................................. 353 Matrix Modulation ................................................................................................................................................ 353 32. Extracellular Matrix Proteins Are Potent Agonists of Human Smooth Muscle Cell Migration ...................................................................................................... 353 Terry L. Kaiura, K. Craig Kent

Introduction ................................................................................................................................................................. 353 Cell Migration .............................................................................................................................................................. 353 Extracellular Matrix Proteins ...................................................................................................................................... 354 Extracellular Matrix Protein and Migration ............................................................................................................... 355 Integrins ........................................................................................................................................................................ 356 Cell Signaling ................................................................................................................................................................ 357 Conclusion .................................................................................................................................................................... 359

Facilitation of Healing .................................................................................................................................. 363 Matrix Modulation ................................................................................................................................................ 363 33. Extracellular Matrix Effect on Endothelial Control of Smooth Muscle Cell Migration and Matrix synthesis ............................................................................... 363 Richard J. Powell

Introduction ................................................................................................................................................................. 363 Model ............................................................................................................................................................................ 364 Effect of Matrix on Endothelial Cell Control of Smooth Muscle Cell Migration .................................................... 364 Matrix Effect on Endothelial Cell Control of Smooth Muscle Cell Matrix Synthesis .............................................. 366

Facilitation of Healing.................................................................................................................................. 371 Cell Entrapment from the Circulation .................................................................................................................. 371 34. Circulating Stem Cells: A Fourth Source for the Endothelialization of Cardiovascular Implants ............................................................................................................................. 371 Willie R. Koen

Introduction ................................................................................................................................................................. 371 The History of Spontaneously Healing Surfaces ........................................................................................................ 371 Properties of Hematopoietic Stem Cells and Progenitor Cells .................................................................................. 372 The Location of Hematopoietic Progenitor Cells ...................................................................................................... 372 Recognition of Stem Cells ............................................................................................................................................ 373 Control of Stem Cells ................................................................................................................................................... 373 Speculative Application ................................................................................................................................................ 375 A Practical Vision for Future Cardiovascular Implants ............................................................................................. 376

Facilitation of Healing.................................................................................................................................. 379 Cell Entrapment from the Circulation .................................................................................................................. 379 35. Surface Population with Blood-Borne Cells ....................................................................................................... 379 William P. Hammond

Background ................................................................................................................................................................... 379 Previous Studies ........................................................................................................................................................... 380 Present Studies .............................................................................................................................................................. 383 Related Findings ........................................................................................................................................................... 383 Implications .................................................................................................................................................................. 384

Facilitation of Healing.................................................................................................................................. 387 Cell Entrapment from the Circulation .................................................................................................................. 387 36. Cellular Population of the Textured-Surface Left Ventricular Assist Devices Leads to Sustained Activation of a Procoagulant and Proinflammatory Systemic Response ...................... 387 Talia B. Spanier, Ann Marie Schmidt, Mehmet C. Oz

Overview ....................................................................................................................................................................... 387 Thrombin Generation and Fibrinolysis ...................................................................................................................... 388 LVAD Surface Cellularization ...................................................................................................................................... 389 Cellular Activation ....................................................................................................................................................... 394 LVAD as Inflammatory/Immune Organ ..................................................................................................................... 395 κB Is a Marker of Cellular Activation in the LVAD Surface Milieu NF-κ and a Target for Anti-Inflammatory Intervention ................................................................................................... 398 Conclusion .................................................................................................................................................................... 400

Facilitation of Healing.................................................................................................................................................. 403 Transdifferentiation .............................................................................................................................................. 403 37. Transdifferentiation and the Vascular Wall ........................................................................................................ 403 William A. Beresford

Endothelial Cells ........................................................................................................................................................... 403 Fibroblasts ..................................................................................................................................................................... 406 Smooth Muscle Cells .................................................................................................................................................... 408 Pericytes ........................................................................................................................................................................ 410 Mast Cells ...................................................................................................................................................................... 410 Macrophages ................................................................................................................................................................. 410 Conclusions .................................................................................................................................................................. 410

Facilitation of Healing.................................................................................................................................. 417 Transdifferentiation .............................................................................................................................................. 417 38. Endothelial Cells Transformed from Fibroblasts During Angiogenesis ........................................................... 417 Takashi Fujiwara, Kazunori Kon

Introduction ................................................................................................................................................................. 417 Angiogenesis in Rabbit Ear Chamber ......................................................................................................................... 417 Discussion ..................................................................................................................................................................... 419 Concluding Remarks .................................................................................................................................................... 422

Facilitation of Healing .................................................................................................................................. 425 Transdifferentiation .............................................................................................................................................. 425 39. Mechanical Forces and Cell Differentiation ....................................................................................................... 425 Ira Mills, Bauer E. Sumpio

Introduction ................................................................................................................................................................. 425 PART III

Biointeractive Prostheses: Complete Healing Engineering Components .....................................................................................................................439 Scaffold Engineering .................................................................................................................................... 441 Structural Designs ................................................................................................................................................. 441 40. An Integral Mathematical Approach to Tissue Engineering of Vascular Grafts ........................................................................................................ 441 Greg R. Starke, A.S. Douglas, D.J. Conway

Introduction ................................................................................................................................................................. 441 Inflammation and the Healing Response .................................................................................................................... 441 The Mechanics of Arteries ........................................................................................................................................... 442 Material Laws for the Native Artery and Graft Materials .......................................................................................... 445 Finite Element Implementation .................................................................................................................................. 452 Conclusions .................................................................................................................................................................. 456

Scaffold Engineering .............................................................................................................................459 Material Aspects .................................................................................................................................................... 459 41. Bioinertness: An Outdated Principle .................................................................................................................. 459 David F. Williams

The Convention of Inertness ....................................................................................................................................... 459 The Impossibility of Inertness ..................................................................................................................................... 460 The Requirement for Controlled Reactivity ............................................................................................................... 461 The Essential Requirement of Reactivity in Tissue Engineering ............................................................................... 462 Conclusions .................................................................................................................................................................. 462

Scaffold Engineering .................................................................................................................................... 463 Material Aspects .................................................................................................................................................... 463 42. Bioinert Biomaterials: Are Their Properties Irreplaceable? ............................................................................... 463 Patrick T. Cahalan

Scaffold Engineering .................................................................................................................................... 469 Material Aspects .................................................................................................................................................... 469 43. Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses ..................................................................................................... 469 Arthur J. Coury

Introduction ................................................................................................................................................................. 469 Design Concept ............................................................................................................................................................ 469 Material Requirements and Limitations ..................................................................................................................... 469 Modes of Polymer Degradation .................................................................................................................................. 470 Polymers with Potential to Serve as Biostable Scaffolds ............................................................................................ 472 Concluding Comments ................................................................................................................................................ 477

Scaffold Engineering .................................................................................................................................... 481 Material Aspects .................................................................................................................................................... 481 44. Biophilic Polymers: What’s on the Horizon? ...................................................................................................... 481 Patrick T. Cahalan

Physical Methods .......................................................................................................................................................... 484 Chemical Methods ........................................................................................................................................................ 486 Summary ....................................................................................................................................................................... 487

Scaffold Engineering .................................................................................................................................... 489 Material Aspects .................................................................................................................................................... 489 45. Bioresorbable Grafts: A Counterintuitive Approach ......................................................................................... 489 David Fox, David A. Vorp, Howard P. Greisler

Overview ....................................................................................................................................................................... 489 Definitions .................................................................................................................................................................... 489 Introduction ................................................................................................................................................................. 489 Single Component Resorbable Grafts ......................................................................................................................... 490 Preventing Aneurysmal Dilatation .............................................................................................................................. 498 Resorbable Outer Wraps .............................................................................................................................................. 500 Biopolymers: Resorbable Prostheses of Biologic Origin ........................................................................................... 500 Conclusion .................................................................................................................................................................... 501

Scaffold Engineering .................................................................................................................................... 505 Material Aspects .................................................................................................................................................... 505 46. Biodegradable Materials ...................................................................................................................................... 505 K.J.L. Burg, S.W. Shalaby

Introduction ................................................................................................................................................................. 505 Criteria for Useful Tissue Engineering Materials ....................................................................................................... 505 Methods of Fabrication of Synthetic Polymeric Materials for Use in Tissue Engineering ...................................... 507 Biodegradable Polymers for Vascular Grafts—Past, Present, and Future ................................................................. 508 Newly Tailored Materials for Tissue Engineering ....................................................................................................... 510 Future Prospectives ...................................................................................................................................................... 510

Scaffold Engineering .................................................................................................................................... 513 Material Aspects .................................................................................................................................................... 513 47. The Influence of Porosity and Surface Roughness on Biocompatibility .......................................................... 513 J.M. Schakenraad, K.H. Lam

Introduction ................................................................................................................................................................. 513 Evaluation Models ........................................................................................................................................................ 514 Interpretation of Results .............................................................................................................................................. 525 Summary and Conclusions .......................................................................................................................................... 528

Scaffold Engineering .................................................................................................................................... 531 Surface Modification ............................................................................................................................................. 531 48. Microgroove Driven Tissue Ingrowth ................................................................................................................ 531 Edwin T. Den Braber, John A. Jansen

Implants, Tissue Engineering, and Biomaterials ........................................................................................................ 531 Surface Micropatterns Manipulating Cellular Behavior ............................................................................................ 532 Changed Cell Shape and Cellular Orientation ........................................................................................................... 533 Cell Attachment on Microtextured Surfaces ............................................................................................................... 534 Other Surface Topography Induced Cell Behavior Alterations In Vitro ................................................................... 538 Microtextured Surfaces In Vivo ................................................................................................................................... 538 The Hypotheses ............................................................................................................................................................ 540 Future Perspectives: Vascular Grafts and Microtextured Surfaces ............................................................................ 540

Scaffold Engineering .................................................................................................................................... 545 Surface Modification ............................................................................................................................................. 545 49. Surface Bonding of Heparin ................................................................................................................................ 545 Patrick T. Cahalan ......................................................................................................................................................................... 545

Scaffold Engineering .................................................................................................................................... 553 Surface Modification ............................................................................................................................................. 553 50. Covalent Grafting of RGD Peptides to Synthetic Surfaces ................................................................................ 553 Nina M.K. Lamba, S.L. Cooper

Introduction ................................................................................................................................................................. 553 Methods to Improve Endothelialization ..................................................................................................................... 553 Immobilization of RGD Peptides ................................................................................................................................ 554 Summary ....................................................................................................................................................................... 558

Matrix Engineering ...................................................................................................................................... 561 51. Hydrogels in Biological Control During Graft Healing ..................................................................................... 561 Jeffrey A. Hubbell

Introduction—Why Hydrogels? .................................................................................................................................. 561 Hydrogel Structure and Synthesis ............................................................................................................................... 562 In Situ Transformations ............................................................................................................................................... 563 Protein and Cell Interactions with Hydrogels ............................................................................................................ 564 Designed Cell Adhesiveness of Hydrogels .................................................................................................................. 566 Nonenzymatic Degradation of Hydrogels .................................................................................................................. 566 Enzymatic Degradation of Hydrogels ......................................................................................................................... 567 Controlled Release ........................................................................................................................................................ 567 Design Principles, Hydrogels for Biological Control During Graft Healing ............................................................ 568

Matrix Engineering ...................................................................................................................................... 571 52. In Vivo Synthesis of Organs Using Collagen-GAG Copolymers ................................................................................................................... 571 Ioannis V. Yannas

Introduction ................................................................................................................................................................. 571 Summary of Evidence for In Vivo Synthesis of Organs ............................................................................................. 571 Methodological Principles of Induced Regeneration ................................................................................................. 572 Models of the Mechanism of Regeneration ................................................................................................................ 573 Conclusions .................................................................................................................................................................. 575

Matrix Engineering ...................................................................................................................................... 577 53. Artificial Extracellular Matrix Proteins for Graft Design ................................................................................. 577 Alyssa Panitch, David A. Tirrell

Introduction ................................................................................................................................................................. 577 Artificial Proteins ......................................................................................................................................................... 578 Artificial ECM Proteins ................................................................................................................................................ 578 Current and Future Directions .................................................................................................................................... 580

Matrix Engineering ...................................................................................................................................... 583 54. Cell-Extracellular Matrix Interactions Relevant to Vascular Tissue Engineering ........................................................................................................................ 583 Stephen P. Massia

Introduction ................................................................................................................................................................. 583 Cell-Extracellular Surface Interactions at the Molecular Level ................................................................................. 584 Biomimetic Cell-Adhesive Biomaterials ..................................................................................................................... 586 Cell-ECM Interactions in Vascular Biology ................................................................................................................ 589 Modulating Cell-Extracellular Interactions for Vascular Tissue Engineering .......................................................... 591 Conclusion .................................................................................................................................................................... 593

Matrix Engineering ...................................................................................................................................... 599 55. Use of Hydroxypropylchitosan Acetate as a Carrier for Growth Factor Release ................................................................................................................................ 599 Keiko Yamamura, Toshitaka Nabeshima, Tsunehisa Sakurai

Introduction ................................................................................................................................................................. 599 Use of HPCHA as a Carrier for bFGF ......................................................................................................................... 600 Influence of HPCHA on Release of bFGF ................................................................................................................... 600 Effect of bFGF-HPCHA on the Endothelialization of Grafts .................................................................................... 602 Concluding Remarks .................................................................................................................................................... 604

Cell Engineering ........................................................................................................................................... 605 56. Transplantation of Transduced Smooth Muscle Cells: A Vehicle For Local and Systemic Gene Therapy ............................................................................................ 605 Randolph L. Geary, Alexander W. Clowes, Monika M. Clowes

Introduction ................................................................................................................................................................. 605 Smooth Muscle Cells as Targets for Gene Transfer .................................................................................................... 606 Vectors for SMC Gene Transfer. .................................................................................................................................. 606 Retroviral Vectors for Ex Vivo Arterial Gene Transfer ............................................................................................... 606 Retroviral Vectors for SMC Gene Transfer in Prosthetic Vascular Grafts ................................................................. 608

Index.............................................................................................................................................................................. 613

EDITORS Peter Zilla, M.D., Ph.D. Cardiovascular Research Unit University of Cape Town Medical School Cape Town, South Africa Chapters 1, 13, 15, 17, 18 Howard P. Greisler, M.D. Department of Vascular Surgery Loyola University Maywood, Illinois, U.S.A. Chapters 26, 28, 45

CONTRIBUTORS James M. Anderson, M.D., Ph.D. Institute of Pathology University Hospital of Cleveland Cleveland, Ohio, U.S.A. Chapter 19 William A. Beresford Department of Anatomy Robert C. Byrd Health Science Center West Virginia University School of Medicine Morgantown, West Virginia, U.S.A. Chapter 37 Karin A. Blumofe Chapter 28 Gary L. Bowlin, Ph.D. Department of Biomedical Engineering University of Akron Akron, Ohio, U.S.A. Chapter 3 Karen J.L. Burg, Ph.D. Department of General Surgery Research Carolinas Medical Center Charlotte, North Carolina, U.S.A. Chapter 46 Patrick T. Cahalan Intelligent Biocides Tewksbury, Massachusetts, U.S.A. Chapters 42, 44, 49

David Cheung University of Southern California School of Medicine Childrens Hospital Los Angeles Division of Surgical Research Los Angeles, California, U.S.A. Chapter 27 Richard R. Clark, M.D. Department of Dermatology State University of New York at Stony Brook Stony Brook, New York, U.S.A. Chapter 31 Alexander W. Clowes, M.D. Department of Surgery Division of Vascular Surgery University of Washington Medical School Seattle, Washington, U.S.A. Chapters 22, 56 Monika M. Clowes Department of Surgery Division of Vascular Surgery University of Washington Medical School Seattle, Washington, U.S.A. Chapter 56 D.J. Conway, M.Sc. Department of Computation & Applied Mechanics University of Cape Town Cape Town, South Africa Chapter 40 S.L. Cooper Chapter 50

Fabiola Cordoba University of Southern California School of Medicine Childrens Hospital Los Angeles Division of Surgical Research Los Angeles, California, U.S.A. Chapter 27 Arthur J. Coury, Ph.D. Focal, Inc. Lexington, Massachusetts, U.S.A. Chapter 43 Lester Davids, M.Sc. Cardiovascular Research Unit University of Cape Town Medical School Cape Town, South Africa Chapter 1

Ruediger C. Braun-Dullaeus Department of Medicine Brigham & Women’s Hospital Harvard Medical School Boston, Massachusetts, U.S.A. Chapter 23 Victor J. Dzau Chairman, Department of Medicine Brigham & Women’s Hospital Harvard Medical School Boston, Massachusetts, U.S.A. Chapter 23 John W Eaton, Ph.D. Department of Pediatrics Baylor College of Medicine Houston, Texas, U.S.A. Chapter 20

Edwin T. Den Braber, Ph.D. Katolieke Universiteit Nijmegen Faculty of Medical Sciences Dental School Nijmegen, The Netherlands Chapter 48

Teddy Fischlein, M.D. Johann Wolfgang Goethe-Universität Frankfurt Klinik für Thorax-, Herz und Gefäßchirurgie Frankfurt, Germany Chapters 8, 18

Manfred Deutsch, M.D. Chirurgischen Abteilung Krankenhaus Wien-Lainz Vienna, Austria Chapters 15, 17, 18

M. Fittkau Department of Thoracic and Cardiovascular Surgery J.W. Goethe University Frankfurt/Main, Germany Chapter 8

Dominic Dodd, M.B.Ch.B., F.R.C.S. Chapter 9

David Fox Chapter 45

Duane L. Donovan, M.D. Summa Health System Akron City Hospital Akron, Ohio, U.S.A. Chapter 10

Takashi Fujiwara, Ph.D. Laboratory Animal Center Ehime University School of Medicine Shingenobu Ehime, Japan Chapter 38

A.S. Douglas Mechanical Engineering Department Johns Hopkins University Baltimore, Maryland, U.S.A. Chapter 40 Terri Dower, B.Sc.(Hons.) Cardiovascular Research Unit University of Cape Town Medical School Cape Town, South Africa Chapter 1

Randolph L. Geary, M.D. Assistant Professor of Surgery & Comparative Medicine Bowman Gray School of Medicine Medical Centre Boulevard Winston-Salem, North Carolina, U.S.A. Chapter 56 Alberto Giudiceandrea Vascular Unit University Department of Surgery Royal Free Hospital and School of Medicine London, U.K. Chapter 2

Werner Müller-Glauser, Ph.D. Department of Surgery Clinic for Cardiovascular Surgery University Hospital Zurich Zürich, Switzerland Chapter 11 Anders Haegerstrand, M.D., Ph.D. Department of Neuroscience Karolinska Institute Stockholm, Sweden Chapter 14 Caroline Gillis-Haegerstrand, M.D., Ph.D. Department of Neuroscience Karolinska Institute Stockholm, Sweden Chapters 14, 16 George Hamilton, F.R.C.S. Royal Free Hospital London, U.K. Chapter 2 William P. Hammond, M.D. The Hope Heart Institute University of Washington School of Medicine Seattle, Washington, U.S.A. Chapter 35 Bo Han University of Southern California School of Medicine Childrens Hospital Los Angeles Division of Surgical Research Los Angeles, California, U.S.A. Chapter 27 Timothy J. Heilizer Chapter 28 Lynn L.H. Huang University of Southern California School of Medicine Childrens Hospital Los Angeles Division of Surgical Research Los Angeles, California, U.S.A. Chapter 27 Jeffrey A. Hubbell, Ph.D. Professor Of Biomedical Engineering Institute of Biomedical Engineering and Department of Materials Swiss Federal Institute of Technology (EHT) Zürich & University of Zürich Zürich, Switzerland Chapter 51

John A. Jansen, DDS., Ph.D. Department of BioMaterials Katolieke Universiteit Nijmegen Faculty of Medical Sciences Dental School Nijmegen, The Netherlands Chapter 48 Terry L. Kaiura Division of Vascular Surgery Beth Israel Hospital Boston, Massachusetts, U.S.A. Chapter 32 K. Craig Kent, M.D. Division of Vascular Surgery New York Hospital/ Cornell University Medical Center New York, New York, U.S.A. Chapter 32 Alisa E. Koch, M.D. Department of Medicine Section of Arthritis and Connective Tissue Diseases Northwestern University Medical School Chicago, Illinois, U.S.A. Chapter 25 Willie R. Koen, M.B.Ch.B., F.C.S. Cardiovascular Research Unit University of Cape Town Medical School Cape Town, South Africa Chapter 34 Kazunori Kon, Ph.D. Laboratory Animal Center School of Medicine Ehime University Shingenobu Ehime, Japan Chapter 38 Pieter Koolwijk, Ph.D. Gaubius Laboraroty Leiden, The Netherlands Chapter 29 K.H. Lam Department of Artificial Organs Division Biomaterials and Biocompatability University of Groningen Groningen, The Netherlands Chapter 47

Nina M.K. Lamba, Ph.D. Department of Chemical Engineering University of Delaware Colburn Lab Newark, Delaware, U.S.A. Chapter 50 Peter I. Lelkes, Ph.D. Director, Laboratory of Cell Biology University of Wisconsin Medical School Milwaukee Clinical Campus Sinai Samaritan Medical Center, Winter Research Milwaukee, Wisconsin, USA Chapter 5 John M. McPherson Biotherapeutic Product Development Department Genzyme Corporation Framingham, Massachusetts, U.S.A. Chapter 31 Michael J. Mann Department of Medicine Brigham & Women’s Hospital Harvard Medical School Boston, Massachusetts, U.S.A. Chapter 23

Victor V. Nikolaychik Laboratory of Cell Biology Department of Medicine University of Wisconsin Medical School Milwaukee Clinical Campus Milwaukee, Wisconsin, U.S.A. Chapter 5 Marcel Nimni University of Southern California School of Medicine Childrens Hospital Los Angeles Division of Surgical Research Los Angeles, California, U.S.A. Chapter 27 Erik L. Owens, M.D. Department of Surgery University of Washington Medical School Division of Vascular Surgery Seattle, Washington, U.S.A. Chapter 22 Mehmet C. Oz, M.D. Columbia University New York, New York, U.S.A. Chapter 36

Stephen P. Massia, Ph.D. Department Chemical, Bio, & Materials Engineering Arizona State University Tempe, Arizona, U.S.A. Chapter 54

Miralem Pasic, M.D., Ph.D. Deutsches Herzzentrum Berlin Klinik für Herz-Thorax und Gefäßchirurgie Berlin, Germany Chapter 11

Sharon O. Meerbaum Summa Health System Akron City Hospital Akron, Ohio, U.S.A. Chapter 10

Alyssa Panitch Department of Polymer Science and Engineering University of Massachusetts Amherst, Massachusetts, U.S.A. Chapter 53

Johann Meinhart, Ph.D. Krankenhaus Wien Lainz Vienna, Austria Chapters 15, 17, 18

Richard J. Powell, M.D. Dartmouth-Hitchcock Medical Center Department of Surgery Lebanon, New Hampshire, U.S.A. Chapter 33

Ira Mills Department of Surgery Yale University School of Medicine New Haven, Connecticut, U.S.A. Chapter 39 Toshitaka Nabeshima Department of Neuropsychopharmacology and Hospital Pharmacy Nagoya University School of Medicine Tsuruma-cho, Showa-ku, Nagoya, Japan Chapter 55

Thomas Schmitz-Rixen Zentrum fuer Chirurgie Johann-Wolfgang Goethe Universität Frankfurt-am-Main, Germany Chapter 2

E. Helene Sage, Ph.D. University of Washington School of Medicine Department of Biological Structure Seattle, Washington, U.S.A. Chapter 30 Tsunehisa Sakurai First Department of Surgery Nagoya University School of Medicine Tsuruma-cho, Showa-ku, Nagoya, Japan Chapter 55 Mark M. Samet Department of Medicine Laboratory of Cell Biology University of Wisconsin Medical School Milwaukee Clinical Campus Milwaukee, Wisconsin, U.S.A. Chapter 5 J.M. Schakenraad Department of Pathology University Hospital Rottendam Dijkzicht, The Netherlands Chapter 47 Ann Marie Schmidt Columbia University New York, New York, U.S.A. Chapter 36 Steven P. Schmidt, PhD Falor Centre for Vascular Studies Akron City Hospital Akron, Ohio, U.S.A. Chapters 3, 10 Alexander M. Seifalian Vascular Unit University Department of Surgery Royal Free Hospital and School of Medicine London, U.K. Chapter 2 S.W. Shalaby Poly-Med, Inc. Anderson, South Carolina, U.S.A. Chapter 46 Paula K. Shireman, M.D. Department of Vascular Surgery Loyola University Maywood, Illinois, U.S.A. Chapters 26, 28

J.Vincent Smyth Department of Vascular Surgery Manchester Royal Infirmary England Chapters 4, 9 Talia B. Spanier, M.D. Columbia University New York, New York, U.S.A. Chapter 36 Greg R. Starke, Ph.D. Department of Computation & Applied Mechanics University of Cape Town Cape Town, South Africa Chapter 40 Bauer E. Sumpio, M.D., Ph.D. Department of Surgery Yale University School of Medicine New Haven, Connecticut, U.S.A. Chapter 39 Zoltan Szekanecz, M.D., Ph.D. Third Department of Medicine University Medical School of Debrecen Hungary Chapter 25 Liping Tang Department of Pediatrics Baylor College of Medicine Houston, Texas, U.S.A. Chapter 20 David A. Tirrell, M.D. Materials Res. Science & Engineering Center University of Massachusetts at Amherst Amherst, Massachusetts, U.S.A. Chapter 53 Marko Turina, M.D. Department of Surgery Clinic for Cardiovascular Surgery University Hospital Zürich Zürich, Switzerland Chapter 11 Victor W.M. Van Hinsbergh, Ph.D. Gaubius Laboratory Leiden, The Netherlands Chapter 29

Robert B. Vernon, Ph.D. Department of Biological Structure University of Washington School of Medicine Seattle, Washington, U.S.A. Chapter 30

David F. Williams, Ph.D., D.Sc. Department of Clinical Engineering Faculty of Medicine Royal Liverpool University Hospital Liverpool, U.K. Chapter 41

Manuela Vici, D.Sc. Institute di Microscopica Elettronica Clinica Universita di Bologna Policlinico S.Orsola Bologna, Italy Chapter 7

Stuart K. Williams, Ph.D. University of Arizona Department of Surgery Tucson, Arizona, U.S.A. Chapters 6, 12

David A. Vorp University of Pittsburgh Pittsburgh, Pennsylvania Chapter 45 Sharon M. Wahl Chief, Oral Infection & Immunity National Institute of Dental Health National Institutes of Health Laboratory of Immunology Bethesda, Maryland, U.S.A. Chapter 21 Michael G. Walker, M.D. Deparment of Vascular Surgery Manchester Royal Infirmary Manchester, U.K. Chapters 4, 9

Keiko Yamamura Department of Hospital Pharmacy Nagoya University School of Medicine Tsuruma-cho, Showa-ku, Nagoya, Japan Chapter 55 Ioannis V. Yannas Department of Mechanical Engineering Massachusetts Institute of Technology Cambridge, Massachusetts, U.S.A. Chapter 52 Ian Zachary The Wolfson Institute for Biomedical Research University College London London, U.K. Chapter 24 John Zagorski Chief, Oral Infection & Immunity National Institute of Dental Health National Institutes of Health Laboratory of Immunology Bethesda, Maryland, U.S.A. Chapter 21

PART I Bio-Inert Prostheses: Insufficient Healing

CHAPTER 1 The Lack of Healing in Conventional Vascular Grafts Lester Davids, Terri Dower, Peter Zilla

Introduction

S

eldom has any prosthetic implant become the nemesis of so many manufacturers as small diameter synthetic vascular grafts. In the era of microsphere-encapsulated cell transplants and human gene therapy it is astonishing that we still do not fully understand why the incorporation of prosthetic cardiovascular grafts does not even remotely resemble healing. Even if the postnatal organism has mostly forgotten the “restitutio ad integrum” healing of the fetal period—the complete regeneration of highly complex tissue with its original cell components—it is nevertheless capable of repairing every nonlethal injury other than chronic infection with a quiescent scar tissue. In the case of synthetic implants, however, the tissue response is a far cry from any concluded repair event. The majority of contemporary prosthetic vascular grafts are not only continually lacking typical components of a healthy native artery such as a midgraft endothelium, contractile arrangements of smooth muscle cells, functionally differentiated elastin formations and other specialized components of a vessel’s extracellular matrix. Synthetic vascular grafts are also lacking the much simpler ability to restore tissue integrity through a quiescent scar formation. Even after prolonged periods of implantation, chronic inflammation dominates the interstices of these prostheses while the luminal patency is endangered by thrombotic surface appositions or the uncontrolled proliferation of a so-called “neointima”. Today’s dilemma is certainly the consequence of a multitude of developments of the past three decades. When modern cardiovascular surgery first set out to replace arteries with synthetic grafts, a graft had simply to act as a nonleaking blood conduit. When it soon became obvious that the surface thrombogenicity of these prostheses was significantly higher than that of native vessels, this complication was seen as a shortcoming of imperfect materials rather than the lack of active physiological functions. This material-centered era only slowly gave way to the realization that the right cells need to be an integral part of a synthetic substitute vessel. However, a full commitment to truly “healing” grafts was further delayed by the industry’s fear of disproportionate delays for new products resulting from a new dimension of complexity. This historical retrospect makes it clear that the most consequential reason for the scanty patchwork of data available today lies in the fundamentally different concepts of graft designs in this earlier era. Instead of a holistic engineering approach which defined all biological components separately before sensibly combining them, grafts were mostly manufactured on the basis of material choices and mechanical strength. Therefore, completed products were investigated with the intention to fulfill regulatory requirements rather than learn about principal mechanisms. As a consequence, a multitude of poorly defined variants such as entirely different synthetics, a wide range of interstitial space dimensions and a variety Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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of animal models make it extremely difficult today to extract significant information from decades of prosthetic graft research. At the same time, biology has only recently provided us with the necessary diagnostic means to analyze tissue reactions with regard to the identification of cell types as well as secretion products. Thus, the emergence of basic biological tools and engineering technologies in the recent past— promising the development of implants which may eventually heal “ad integrum”—does not build on a comprehensive knowledge regarding the shortcomings of previous concepts. Since the retrospective deduction of biological principles from data which had mostly been obtained with a different intention will not be able to create a full understanding of the apparent absence of physiological healing patterns in prosthetic vascular grafts, a systematic experimental approach will need to supplement the erratic experience of the past thirty years. For any eventual conclusions preceding serious tissue engineering approaches of the future, well planned, biologically oriented in vivo and in vitro experiments will need to be performed. In this chapter we will nevertheless attempt to retrospectively identify certain criteria which may have influenced the mitigated healing response of prosthetic vascular grafts in the past. In order to avoid dealing with an infinite number of variables, we have concentrated our efforts on those two basic graft types which have been dominating the field of synthetic vascular prostheses for the past decades: Dacron and expanded polytetrafluoroethylene.

Midgraft Healing No other aspect of prosthetic graft research missed the point as fundamentally as did midgraft healing. In spite of a unanimous agreement with regard to the limits of transanastomotic tissue outgrowth in man, the vast majorFig. 1.1. Schematic diagram depicting the histological summary of events over time for different experimental models for low porosity (< 30 µm) ePTFE. Detailed histological descriptions can be found in the text.

Tissue Engineering of Prosthetic Vascular Grafts

ity of studies of the last three decades focused unintentionally on exactly this type of healing. By choosing animal models which far exceed the human ability of transanastomotic healing, as well as graft lengths which were in average one tenth of those clinically used, midgraft healing under experimental conditions occurred predominantly through anastomotic outgrowth rather than transmural ingrowth. Considering the fact, however, that prosthetic arterial graft endothelialization in humans—mainly restricted to the anastomotic region—often involves less than one thirtieth of the entire surface, facilitated anastomotic outgrowth will most likely continue to play a limited role. Therefore, in spite of the countless studies which investigated prosthetic arterial grafts, a description of midgraft healing—clearly distinguishable from transanastomotic events—is the exception rather than the rule. Nevertheless, by putting the sporadic observations of all these studies into a framework, a surprisingly consistent pattern of healing became apparent. A further confusion arising from the multitude of animal models used also proved to be less of a problem than expected. As with any other aspect of graft healing, by and large midgraft healing also confirmed the previous belief that it is not the healing pattern itself but its time course which distinguishes the various animal models from each other. However, there are healing aspects like the impenetrable inner fibrin capsule of Dacron grafts in humans which may well be based on principal differences between the species rather than simply on a protracted time course of common events. Nevertheless, these principal differences may again reflect differences in the time course of common sequences, as certain biological phenomena may only occur at a certain point in time. Therefore, a solution to graft healing could well lie in the facilitation of the ideal timing of crucial biological events.

The Lack of Healing in Conventional Vascular Grafts

ePTFE Grafts In spite of the wide range of internodal distances which were experimentally evaluated and the distinct characteristics of each of those grafts, a certain cut off point emerged beyond which tissue ingrowth has seemed to be principally possible. This cut off point was identified to lie somewhere between 30 µm and 45 µm of internodal distance. Therefore, it became customary to refer to grafts with an internodal distance of less than or equal to 30 µm as low porosity ePTFE grafts and to those of 45 µm or more as high porosity ePTFE grafts. Low Porosity ePTFE Grafts (< 30 mm Internodal Distance) In contrast to other prosthetic grafts, low porosity ePTFE hardly shows any difference between animal models with regard to the time course of midgraft healing.

5

During the initial 2 weeks of implantation, a finefibrillar layer of a fibrin coagulum of approximately 15 µm thickness covers the graft surface.1,2 During the following weeks the thickness of this thrombus does not further increase,3,4 but its composition becomes more variable, reaching from acellular loosely3,5 (Fig. 1.2) or highly compressed fibrin4,23 to platelet carpets (Fig. 1.3) with interspersed granulocytes.1,5 Interestingly, these dense layers of sometimes almost pure platelets often show a morphology of low grade activation (Fig. 1.3) or even nonactivation.5 In spite of this lack of morphological signs of activation, these platelets almost seem to melt into a dense, amorphous matrix.5 During the subsequent months4,6,7 the blood surface still remains covered by either unorganized, compacted fibrin or a more amorphous platelet rich material, but the thickness of this layer increases to about 80-290 µm.4,5,8 Even after prolonged observation periods of up to 3 years, midgrafts lack all forms Fig. 1.2. Surface fibrin on ePTFE (30 µm internodal distance) after 4 weeks of implantation in the chacma baboon model. Loosely covering fibrin meshwork on a dense carpet of platelets adheres to a compact, more amorphous thrombus in the depth.

Fig. 1.3. Surface of ePTFE graft (30 µm internodal distance) after 4 weeks of implantation in the chacma baboon. One can recognize the ePTFE structure in the left lower corner. A relatively thin layer of amorphous thrombus covers the synthetic surface. Densely adherent platelets which are mostly discoid in shape seem to melt into the amorphous protein layer.

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of cellular coverage.6,9-11 However, anecdotal descriptions of multilayered cell islands in the central segments of grafts1,5 add to the decades old controversy concerning the origin of intimal cells.12 Similar to surface healing, transmural tissue ingrowth remains incomplete. Until week 2 most of the interstices of the grafts are devoid of any recognizable cellular material,1 with only scanty macrophages beginning to migrate into the micropores.13 After 3 weeks the fibrin matrix in the graft wall has slowly become populated with inflammatory cells,2 but there is hardly any connective tissue ingrowth into the graft.13-17 In the chacma baboon model we find a typical triple layer picture, with macrophages invading from the blood surface and the adventitia into an otherwise almost acellular central fibrin matrix (Fig. 1.4). Apart from scanty

Fig. 1.4. 30 µm ePTFE graft after 6 weeks in the chacma baboon. One can clearly recognize the fine porous wrap at the outer surface. A layer of foreign body giant cells (FBGC) demarcates the ePTFE from an otherwise wellhealed surrounding tissue. Loose and scanty cell infiltrates from the blood surface and the adventitial side simultaneously progress towards the acellular central fibrin matrix, filling the interstices.

Fig. 1.5. Schematic diagram depicting the histological summary of events over time for different experimental models for high porosity (>45 µm) ePTFE. Detailed histological descriptions can be found in the text.

Tissue Engineering of Prosthetic Vascular Grafts

multinucleated giant cells13,18,19 and macrophages8-20 demarcating the outer graft surface against a well developed fibroelastic perigraft tissue,1,7 no persistent inflammatory reaction is found around the graft.7 However, there is a distinct difference between wrap-reinforced and nonreinforced grafts with regard to the tissue interaction with the outside environment. In nonwrapped grafts, relatively few foreign body giant cells (FBGC) adhere to the PTFE nodes on the outside surface, while single macrophages have infiltrated the interstices. In wrapped grafts, the narrow structure of the wrap limits the infiltration of the interstices with cells. Therefore, the almost acellular fibrin matrix usually found in the center of the prosthetic wall extends all the way to the outside surface. This outside surface, in return, borders upon a distinct layer of macrophages and FBGC (Fig. 1.4). This indi-

The Lack of Healing in Conventional Vascular Grafts

cates once more that the formation of FBGC may be significantly influenced by surface structures, because the material is PTFE in both instances. In the case of the ePTFE wrap it is difficult to say whether the higher density of FBGC is a result of the finer porosity and structure of the PTFE or the overall result of an impenetrable barrier plane compared to the relatively open spaced single nodes in unwrapped ePTFE. Eventually, after 6 weeks of implantation, connective tissue cells occasionally begin to grow into the interstices of unwrapped prosthesis from outside.13,14,17,21-23 Six months after implantation, however, connective tissue ingrowth from outside still remains moderate and mostly limited to the outer part of the graft wall,18-20,24-26 while the majority of interstitial graft spaces continue to be primarily occupied by fibrin and a mild inflammatory infiltrate in the proximity of the surfaces.25 In contrast to the early days of implantation, however, these cells lie in a very fine fibrillar extracellular meshwork which has replaced the solid acellular fibrin. If present, this scanty and loose connective tissue contains hardly any capillaries.113,25 High Porosity ePTFE Grafts ( > 45 µm Internodal Distance) In contrast to low porosity ePTFE grafts, there is a distinct difference in high porosity ePTFE grafts with regard to the time course of healing events in humans and different animal models. Initially, the blood surface resembles that of low porosity ePTFE grafts. It is covered by a thin fibrin layer which over time develops into a variable coagulum of fibrin, platelets and erythrocytes.27 In dogs and other more senescent animal models, this stage lacking endothelial coverage persists well into the sixth week of implantation.27-29 In the senescent chacma baboon, for instance, we find a six week pannus ingrowth of only 7.8 ± 3.5 mm reaching onto otherwise nonendothelialized grafts (Fig. 1.6). The surface of these grafts is covered either by a relatively thin layer of smooth fibrin containing white blood cells and a few platelets or by

7

a more amorphous protein matrix densely covered by platelets and a few white blood cells (Fig. 1.7). Overall, the thickness of the surface thrombus is moderate, because one can recognize the nodal ePTFE structure throughout the graft (Fig. 1.8). Histologically, a homogenous fibrin layer covered by a carpet of platelets enshrouds the ePTFE surfaces. Except for very few occasional polymorphnuclear granulocytes, this fibrin layer is practically acellular (Fig. 1.7). The presence of a layer of macrophages underneath this fibrin matrix—which is demarcated for the depth of the graft by a zone of equally acellular fibrin—indicates that the surface fibrin has previously been permeable to inflammatory cells. In contrast to the senescent chacma baboon, high porosity ePTFE grafts lead to early and spontaneous endothelialization in juvenile recipients such as young yellow baboons. Patches of endothelial cells30 and capillary orifices—approximately 100-500 µm apart31—appear on the blood surface as early as 1-2 weeks after implantation, leading to confluence shortly thereafter.30-35 Since the endothelium soon rests on a well developed layer of actin positive cells3-33,35 which contains only a few macrophages,36 the fibrin matrix which initially covers the inner surface and provides the outgrowth substrate for the endothelium represents a very transient matrix under these circumstances. These actin positive cells appearing underneath the endothelium9,33 also exhibit the ultrastructural characteristics of arterial smooth muscle cells.30 Although proliferating smooth muscle cells resemble wound fibroblasts, there is currently more evidence against fibroblasts trans-differentiating into smooth muscle cells19,37-40 than there is for it. Therefore, it is likely that these cells are derived from pericytes accompanying EC.9,33 Over time the endothelium and its subendothelial tissue layer develop into a stable neointima20,31,35,41,42 with a higher density of actin positive cells adjacent to the endothelium than to the graft surface.20,26,31,40,43-49 This intimal thickening is evenly distributed along the entire graft surface and not confined to the anastomotic region as in low porosity ePTFE grafts.30 It is

Fig. 1.6. Pannus outgrowth onto 60 µm ePTFE after 6 weeks of implantation in the chacma baboon. One can clearly see the anastomotic outgrowth of the pannus and the endothelium tapering off onto an otherwise nonendothelialized ePTFE surface.

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Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.7. Platelet-rich surface coverage of the midsection of a 60 µm ePTFE graft after 6 weeks of implantation in the chacma baboon. One can faintly recognize the ePTFE surface in the central upper part of the picture. The dense carpet of platelets is again dominated by discoid shaped thrombocytes which do not resemble the typical morphology of aggregated platelets.

Fig. 1.8. 60 µm ePTFE after 6 weeks of implantation the chacma baboon. The surface thrombus is delicate and still allows the underlying PTFE structure to be recognized.

particularly noteworthy that these smooth muscle cells only started to appear in the intima and began to proliferate after endothelial cells had covered the luminal surface.30 With SMCs multiplying in the inner one third of the intima adjacent to an endothelium33 and without the concomitant presence of platelets33 the situation is the opposite of traumatized arteries. Therefore, it is conceivable that endothelial cells themselves produce growth promoting substances for smooth muscle cells. This is further supported by the fact that after 3 months, SMCs are predominantly found in the subendothelial intima relatively distant from the macrophages in the interstices.33 Apart from hinting at the involvement of endothelial cells in intimal smooth muscle cell proliferation, this observation of an increasing smooth muscle cell proliferation with increasing distance from the macrophages is inconsistent with the hypothesis that macrophages are an important source of SMC mitogens,134,189 unless an inverse gradient applies in which low concentrations of

macrophage products stimulate SMC proliferation and high ones are inhibitory.69 Eventually the discovery of PDGF-A mRNA in the overlying endothelial cells resolved the issue of the source of SMC mitogens.114b This cytokine is not only one of the strongest currently known proliferative agents for smooth muscle cells, but was previously already shown in vitro to be secreted by endothelial cells vectorially into the abluminal basal compartment.228 Its strong presence also explains the 5-100 times higher proliferative activity of SMCs and the 10-100 times higher one of endothelial cells50,51 in grafts than in normal arteries. Nevertheless, the presence of higher levels of PDGF and the proof of a higher mitotic activity does not indicate whether this is an adaptive response or a primary phenomenon. On the one hand one could argue that a high cycling rate may simply be the consequence of an upregulated PDGF production. On the other hand, however, it is also quite likely that a poor anchorage of endothelial cells at this stage leads to mechanical detachment

The Lack of Healing in Conventional Vascular Grafts

and this in turn upregulates PDGF production as a reparative response. The latter explanation is supported by the morphological appearance of endothelial cells. Typical for endothelial surfaces affected by a distinct cell loss,5 these endothelial cells lack the regular spindle shaped pattern30 and are larger than those in arteries.30 This sequence of a higher cell loss preceding a higher mitotic activity, rather than a higher cycling rate, being the reason for higher endothelial cell shedding would coincide with the finding for in vitro endothelialized grafts that a distinct degree of endothelial cell detachment occurs during early implantation in spite of continual endothelial integrity.5 However, since the overall thickness of this neointima may also represent an adaptation attempt of the tissue to flow conditions in an otherwise rigid graft, the PDGF upregulation in the surface endothelium could as a third explanation also be a natural mechanism of macrovascular endothelial cell response to environmental changes such as flow. This would explain why a many-fold thicker neointima is found under low flow conditions than under high flow conditions.54 When surface healing occurs so early and independently of transanastomotic tissue ingrowth, the cell source for it can only be perigraft tissue which reaches the blood surface through apt transmural ingrowth. However, since complete transmural ingrowth does not readily occur in other graft types, it is particularly interesting to see whether it differs substantially in high porosity ePTFE grafts compared with other prostheses. Comparably to other grafts, the basically acellular earliest interstitial fibrin matrix gets progressively populated with macrophages and polymorphnuclear leukocytes28,33 during the initial period of graft implantation followed by capillary ingrowth. In the juvenile yellow baboon, these capillaries reach through the entire wall as early as after 2 weeks of implantation. Although fibroblasts from the perigraft region soon follow the capillaries,16 the majority of the graft interstices still remain populated with macrophages.33 In more senescent animal models like the dog or the chacma

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baboon, it takes 2-5 weeks and longer until sprouting capillaries begin to reach into the macrophage dominated outer one third to one half of the graft wall27-29 (Fig. 1.9). Distinctly different from Dacron grafts, these microvessels resemble mature arterioles which are small in diameter (23.06 ± 13.11 µm; personal observation) and regularly accompanied by at least one layer of smooth muscle cells. However, in spite of this initial vascularization, the histological picture of these implants remains dominated by a triple layer appearance even after 6 weeks of implantation: While the blood and the adventitial part of the prostheses is infiltrated by inflammatory cells, the central zone of the prosthetic wall often remains filled with an almost acellular amorphous matrix (Fig. 1.10). Within this triple layered structure the cellularity of the outer third is regularly higher than that of the inner third of the graft, although it does not change much over time. Ham 56 positive macrophages continue to dominate over connective tissue cells,55 while foreign body giant cells are conspicuously absent.56 In spite of the scanty presence of connective tissue, the homogenous fibrin matrix initially filling the interstices gradually gets replaced by a loose transparent extracellular meshwork wherever inflammatory cell infiltrates reach into the interstices (Fig. 1.11). On the outside, high porosity ePTFE grafts are embedded into a moderately developed mature fibrous tissue which contains hardly any inflammatory cells. Only the direct interface between adventitial tissue and graft occasionally shows a few single FBGCs.

Dacron Grafts In Dacron grafts the issue of porosity is slightly complicated through the optional combination of the grafts with a velour surface which differs from the basic texture. Low Porosity Dacron Grafts (Woven) Immediately after implantation a thin layer of fibrin, erythrocytes, white blood cells and platelets is deposited on the blood surface of the prosthesis.40,57 During the first few

Fig. 1.9. Immunohistochemical staining of α-actin in a 60 µm ePTFE graft after 6 weeks of implantation (chacma baboon). It is obvious that those vessels which were capable of penetrating into the outer half of the graft all contain at least one layer of smooth muscle cells.

10

hours to days this thrombus layer slowly starts thickening until it reaches a stable equilibrium of compacted fibrin.45 In dogs this stage is reached within 6 months,45 whereas in humans it is observed after 1 1/2 years.223 Exceptionally, fibrous tissue begins to organize the basis of the fibrin capsule in a few scanty areas of the graft.40,45 This process is observed after 4-8 weeks in dogs40,45 and after 1 1/2 years in humans.58 Those few areas of a pseudoneointima where tissue reaches the blood surface40,58,59 have a thickness which is comparable to that of the surrounding compacted fibrin (400-500 µ m). 58 In humans these rare and localized neointimal spots extend over less than 1 mm and are collagen rich, with increasing cellularity towards the graft surface.47 The situation is similar with regard to EC. In areas which represent true midgraft regions—unaffected by anastomotic ingrowth—no endothelial cells are found within the first 1-2 years.40,58,59 After prolonged implantation periods of up to 11 years—which are naturally restricted to humans—small islands of endothelial cells may occasionally

Fig. 1.10. 60 µm ePTFE graft without external wrap reinforcement. There is hardly any inflammatory tissue demarcating the graft against the surrounding, well-healed fibrous adventitial tissue. After 6 weeks the central zone of the interstitial spaces is still filled by an acellular fibrin matrix. The outer and inner third of the graft shows a loose infiltration with cells which previously degraded the acellular fibrin matrix. The blood surface is covered by a thin irregular thrombus which is practically acellular, not even containing inflammatory cells.

Fig. 1.11. Knitted Dacron graft (Vascutec) after 6 weeks of iliac interposition in the chacma baboon. Typically, the Dacron structure is packed with foreign body giant cells and demarcated from the surrounding adventitia by granulation tissue which contains hugely dilated capillary sinuses.

Tissue Engineering of Prosthetic Vascular Grafts

appear.58,59 These endothelial islands always rest on the compacted fibrin rather than on the scanty islands of connective tissue. Underneath this apparently impenetrable fibrin layer, healing is most likely to take place; at least with regard to perigraft tissue ingrowth, the initial events of graft incorporation are similar to ePTFE. A fibrin matrix fills the narrow prosthetic interstices within minutes of implantation,57 and organizing tissue begins to invade the fibrin layer surrounding the graft to form the outer fibrous tissue capsule.45,57 In the interface between this fibrous tissue and the Dacron yarns a variable degree of foreign body giant cells begins to build up.57 Subsequently, a few capillaries and fibroblasts start growing into the tight interstitial spaces of the woven Dacron.40,57 This process is quite variable in its extent and time course. In dogs it can be seen as early as 2-3 weeks after implantation,40 whereas in humans it may be observed after 5 months11 but may also be absent after up to 18 years.47,58,59 However, even if tissue does occasionally grow through the

The Lack of Healing in Conventional Vascular Grafts

interstices of the prostheses, it does not break through the compacted fibrin of the inner capsule to reach the blood surface43 even after decades of implantation. High Porosity Dacron Grafts (Knitted) There is a distinct difference regarding the healing response to variations in porosity between Dacron grafts and ePTFE grafts. In ePTFE grafts, the fibrin matrix may get rapidly replaced by inflammatory and connective tissue if the porosity is increased. In Dacron grafts, the mighty inner fibrin capsule mostly persists in both high and low porosity grafts although interstitial healing is more accelerated in the higher porosity prostheses. In knitted dacron prostheses the initial surface net of fibrin, white blood cells, erythrocytes and platelets48 increases during the first week of implantation to a thickness of 100-120 µm.44,60 During the subsequent 5 weeks of implantation it appears as if the inner fibrin capsule has not yet reached an equilibrium and therefore holds a wide range of thickness: 100-500 µm in humans,60 30-210 µm in dogs2,61-66 and 290-300 µm in the yellow baboon.35,91 After a subsequent further insignificant increase,60-62 the thickness of the internal capsule eventually reaches an equilibrium before sixth months of implantation.61,62 Although a superficial screening of the literature suggests that it is only a matter of time until this entire layer of compacted fibrin gets organized by ingrowing tissue, deeper analysis makes it likely that this may well be a misinterpretation based on the confusion of midgraft healing with anastomotic outgrowth. It is true that the vast majority of studies describe a mature endothelium and a well developed intimal smooth muscle layer, resting directly on the graft surface,2,67-69 but the graft length and implantation period in these studies make it almost certain that this mature neointimal tissue represents nothing but extended anastomotic outgrowth.2,22,32,35,44,68-70 This suspicion is confirmed by descriptions of true midgraft re-

11

gions in particularly long grafts in which the inner fibrin capsule appears to be both impenetrable for the ingrowing interstitial graft tissue43,44,48,59,60,71,72 and nonendothelialized. 60,72,73 This lack of surface healing occurs43,44,48,59,60,71,72,74,75 although transmural tissue reaches the compacted fibrin of the blood surface within 3-4 weeks in the calf,60 the yellow baboon35 and the dog,2,76 and 3-6 months in humans.48,60 However, in spite of this dominant healing deficit, complete tissue replacement of the inner fibrin capsule may sporadically occur in the absence of a surface endothelium and with a poorly differentiated fibroblast matrix rather than with smooth muscle cells.61 Occasionally, even complete endothelialization of very long prostheses is seen46,77 comparable to those in highly porous ePTFE grafts. In the yellow baboon, for instance, knitted Dacron grafts may form complete endothelial linings, although not as readily and consistently as in 60 µm ePTFE.32,35,78,79 They initially show small islands of endothelium after 2 weeks, and in a minority of implants even confluence after 4 weeks.32 However, in contrast to the case of high porosity ePTFE grafts, it remains unclear whether this endothelium is derived from transmural tissue ingrowth, facilitated transanastomotic outgrowth or fall-out healing.12,80 The latter is a phenomenon predominantly observed in Dacron grafts where—apparently independently of transmural tissue ingrowth and with some delay—endothelial islands emerge on the surface of the compacted fibrin. These islands are well separated from the transmurally ingrowing fibrous tissue by a distinct layer of compacted fibrin.43,48,59,60,71 They either rest directly on the fibrin43,59,72 or on a few layers of actin positive cells with no other tissue connection,59,71 an observation which can be made in dogs after 2-6 months72 and in humans perhaps in every fourth graft after 1-11 years.43,59,71 However, although sporadic endothelialization occasionally occurs in limited areas of the Fig. 1.12. Schematic diagram depicting the histological summary of events over time for different experimental models for woven Dacron. Detailed histological descriptions can be found in the text.

12

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.13. Schematic diagram depicting the histological summary of events over time for different experimental models for knitted Dacron. Detailed histological descriptions can be found in the text.

grafts, the majority of the surface remains covered by fibrin only. If one focuses on the transmural tissue ingrowth itself, it also begins with the infiltration of the surrounding outer fibrin matrix by granulation tissue.2,48 Initially it is primarily macrophages which begin to invade the interstitial fibrin,14 populating the entire depth of the graft at the end of the first month of implantation.60 Early giant cell formation surrounding the yarns and—if present—the velour fibrils commences after 7 days and increases, predominantly in the outer half of the graft wall,62 during the following months.61,62 The subsequent ingrowth of blood vessels and fibroblasts varies not only between the different animal models but also within one and the same recipient species. In dogs for instance, the interstitial tissue ingrowth can either reach the inner fibrin capsule as early as 2-4 weeks after implantation2,61,63,76 or be absent as late as after 2 months.72,81 Similarly, clinical implants in humans show all stages of ingrowth for up to three months.179 From then onwards, ingrowing adventitial tissue reaches the internal fibrous capsule without exception in all animal species.2,11,43,59,60,71,81,82 This tissue contains a moderate number of capillaries, fibroblasts and collagen2,11,43,59,71,82 and eventually extends along the inside surface of the graft.60,62,71 However, a clear, layered structure of tissue elements still dominates the histological picture. The Dacron fibers appear to demarcate themselves expressively against the surrounding tissue by a particularly distinct accumulation of foreign body giant cells (FBGC). Between this outer layer of the prosthetic wall which is packed with FBGC and the surrounding connective tissue capsule with its longitudinally aligned collagen bundles,59,60,81 hugely dilated capillary sinuses are usually present during the early weeks of implantation. These sinuses typically consist of a single endothelial cell layer which contains no smooth muscle cells or cell types other than macrophages and FBGC, which snuggle against

its basal lamina (Fig. 1.11). Those few capillaries which sometimes reach through the entire graft wall and reach the inner capsule also remain free of smooth muscle cells. However, in areas where transanastomotic ingrowth has resulted in a mighty neointima covering the surface, one occasionally finds true multilayered vessels with smooth muscle cells underneath the endothelium. In summary, the porosity of knitted, and to some extent even woven, Dacron prostheses eventually allows a certain degree of tissue ingrowth from outside. At the same time, however, it seems that the compacted inner fibrin capsule on the blood surface of Dacron grafts represents an ingrowth barrier for transmural connective tissue regardless of whether the grafts are knitted or woven. Although this barrier may eventually be overcome by ingrowth of undifferentiated fibrous tissue, capillaries remain unconnected to the often poorly endothelialized blood surface.75 The peculiar coincidence of zones which are densely packed with foreign body giant cells and hugely dilated capillary sinuses in the vicinity is particularly obvious in Dacron prostheses. Moreover, it is also interesting that the fibrin coverage on Dacron grafts seems to have a higher propensity to capture blood-borne cells than that on ePTFE grafts.80 In this context it is also noteworthy that the resulting isolated endothelial islands often rest on equally isolated smooth muscle layers.43,59 However, in spite of the scientific excitement arising from this observation, the rare and late occurrence of the phenomenon makes it the least typical healing characteristic of Dacron grafts. High Porosity Dacron Velour The rationale behind a filamentous coverage of Dacron fabric was the attempt to improve the healing between perigraft tissue and the prosthesis. Initially only used as an external cover,60 it was soon also applied as an internal cover.83 Interestingly, it was thought that a finer and more

The Lack of Healing in Conventional Vascular Grafts

even porosity would cover the coarser structure of the knitted fabric underneath and thus achieve facilitated healing. As is obvious from our porosity analysis (see “Porosity”) the opposite is true, and velour materials actually provide significantly wider ingrowth spaces than knitted or woven structures. This wider porosity does result in a firmer tissue integration on the outside,15,84 but reduces platelet survival and leads to less pronounced pseudointimal development on the inside.85 In order to study the healing pattern of entire velour structures, we have fabricated graft tubes out of pure velour material and studied the tissue reaction in the chacma baboon. Similarly to previous results,86 the transanastomotic outgrowth of a neointima is significantly less distinct than in nonveloured knitted Dacron, even after weeks of implantation. The fibrin matrix itself often appears looser than on ePTFE and, moreover, often covered with relatively densely adherent macrophages (Fig. 1.14). The loose fibrin often

13

contains white blood cells in deeper layers which one can still recognize from the surface (Fig. 1.15). Nevertheless, one also finds areas of densely packed fibrin, covered with morphologically mostly unactivated platelets and white cells comparable to ePTFE (Fig. 1.16). Very rarely does one see vessel openings on the surface (Fig. 1.17). The relatively freely lying single fibers of the material are completely engulfed by conglomerates of foreign body giant cells, whereas the interstices are dominated by a loose fibrin matrix containing lymphocytes, polymorphonuclear granulocytes, macrophages and FBGCs. In spite of the wide open porosity, hardly any fibroblasts or collagen are found in the interstices. The huge and dilated capillary sinuses, which were all localized between the outer layer of the graft and the surrounding tissue in knitted prostheses, are equally present but scattered throughout the meshwork of the grafts. These often enormously dilated capillary sacs are all factor VIII positive, and

Fig. 1.14. Blood surface of the midgraft section of a Dacron velour tube (6 weeks, chacma baboon). Typically, large areas of the fibrin surface are covered with dense layers of spreading macrophages.

Fig. 1.15. Same graft as Figure 1.10, but showing a more loose meshwork of fibrin with inflammatory cells lying on the surface and in the depth of the structure. One can still recognize those white blood cells which are lying close to the surface.

14

Fig. 1.16. Midgraft section of a Dacron velour tube covered by an amorphous protein layer with an adherent carpet of disk shaped platelets. Occasionally, white blood cells lie in lacunar spaces which appear to have been excavated from the fibrin layer.

Fig. 1.17. Midgraft section of a Dacron velour tube after 6 weeks of implantation in the chacma baboon. One very rarely sees capillary openings surrounded by small islands of endothelial cells, which grow onto the otherwise endothelial-free compacted fibrin surface.

Fig. 1.18. Histological cross-section of a Dacron velour tube after 6 weeks of implantation in the chacma baboon. The loose structure allows capillary ingrowth at all levels. As a consequence the dilated capillary sinuses—which are otherwise only found in the outer half of knitted prostheses—are scattered throughout the graft wall. The capillary sinuses themselves consist of factor viii positive endothelial cell monolayers resting on basement membranes surrounded purely by inflammatory cells, which are predominantly foreign body giant cells.

Tissue Engineering of Prosthetic Vascular Grafts

The Lack of Healing in Conventional Vascular Grafts

equally as devoid of smooth muscle cells as knitted prostheses (Fig. 1.18). Occasionally, the simultaneous infiltration mode from inside and outside can still be detected after 6 weeks, with a central cell free zone of fibrin wedged between the inner and outer graft thirds, the outer being dominated by foreign body giant cells. These observations indicate that in spite of wide open interstitial spaces, there seems to be a distinct inhibition for connective tissue ingrowth in Dacron scaffolds. It is even more obvious from velour explants than from the narrower structures of knitted Dacron that capillarization is rather overshooting than underdeveloped. These widely dilated endothelial sacs, however, neither reach the blood surface nor attract concomitantly ingrowing smooth muscle cells.

Transanastomotic Healing The assessment of anastomotic tissue ingrowth is another ambiguous task within the attempt to interpret the healing of prosthetic vascular grafts. Although surface endothelialization—which under experimental circumstances overwhelmingly happens through transanastomotic ingrowth—was indirectly the focus of almost every single study in the past, there is hardly any other aspect of prosthetic graft healing which is so vaguely described. When all studies included in this review are taken into account, the average graft length is 10.8 ± 2.6 cm. This value includes the 11% of grafts which were longer than 10 cm. If those grafts—which were substantially longer than the others (36.6 ± 2.4 cm)—are excluded, the mean length of the remaining 89% of prostheses is only 5.5 ± 1.2 cm. With a median implantation period of 91.8 ± 17.3 days in Dacron grafts and 60.0 ± 9.36 days in PTFE grafts, it does not come as a surprise that transanastomotic ingrowth was frequently completed before the prostheses were retrieved. This often makes it impossible to determine retrospectively the point in time when the surface coverage through transanastomotic tissue ingrowth was completed. To create further uncertainty, measurements of tissue ingrowth vary widely amongst those studies where surface endothelialization was not completed prior to explantation, and which could thus be admitted to a comparative ingrowth assessment. Another point of confusion relates to the terms “neointima” and “pannus”. As we have argued under the topic “midgraft healing”, prior to the use of high porosity ePTFE grafts neointimal tissue formation was mostly restricted to transanastomotic ingrowth, even if many of the investigators were not aware of it. As opposed to the mostly fibrinous “pseudoneointima” of midgraft regions, “neointima” could almost synonymously be used as the term for transanastomotically ingrowing tissue covering the blood surface of the graft. However, with truly transmurally derived neointimas observed in certain contemporary prostheses, it seems necessary to restrict the term “neointima” to surface tissue which results from transmural midgraft healing and to keep describing the anastomotic tissue ingrowth as “pannus”. As necessary as this restriction of the term “neointima” is, it eliminates further bits of assumed clarity. Since the majority of studies describe prostheses in which the anastomotic pannus has already reached the midgraft

15

region, the widely used term “neointima” in these studies actually refers to anastomotic pannus tissue. However, in spite of the poorly defined criteria with regard to anastomotic tissue ingrowth and the scanty data available for some of these criteria, certain trends concerning animal models, material and porosity still become apparent.

Pannus Tissue Within less than 2 days of implantation, a straightening and disruption of the internal elastic membrane of the adjacent native artery occurs.87 Simultaneously, the smooth muscle cells of the inner one-third of the media begin to proliferate87 and migrate through the ruptured internal elastic membrane into the intima. It is primarily this hyperplastic intimal tissue which—together with the endothelium— migrates and proliferates onto the graft surface, thereby slowly replacing the inner fibrinous capsule from the anastomosis towards the midgraft. Typically, endothelial cells precede this outgrowth,19,28,88 followed by connective tissue cells. This means that the foremost edge of the outgrowing endothelium initially rests on the fibrin matrix of the inner capsule6,28,88 before a few layers of smooth muscle cells appear underneath.14,25,27,13 This edge of outgrowing endothelium often does not resemble a solidly migrating formation of endothelial cells, but rather deeply meandered, tattered and irregularly shaped islands and tongues of endothelium (Fig. 1.19). Very often, the foremost edge consists of loosely lying endothelial cells, surrounded by platelet-covered fibrin (Fig. 1.20). Eventually, a multilayered connective tissue, which increases in thickness over time, slowly builds up between the graft and the blood surface.19 This tissue does not consist exclusively of secretory smooth muscle cells69 but also of fibroblasts, collagen and a few macrophages.30,40,52,55,61,62,68 After a few months a layered structure develops with a hypercellular myofibroblast-dominated luminal aspect and a fibro-collagenous and glucosaminoglycan-rich aspect near the graft surface.68,89 However, even at that stage this subendothelial tissue lacks features of native arteries like elastic fibers between the smooth muscle cell layers20 and an internal elastic membrane69 separating the endothelium from the bulk of smooth muscle cells. Altogether, the undifferentiated pannus near the anastomosis of a prosthetic graft has not much in common with the organized intimal structures of a native vessel. Thus, transanastomotic tissue outgrowth remains a far cry from a self-terminating healing process leading to true neo-artery structures on the inside of a prosthetic graft. In spite of the undifferentiated hyperplastic nature of this pannus tissue, however, there are various circumstances which significantly affect at least its tissue composition and thickness. Anatomical position, for instance,72,81 as well as the difference between distal and proximal anastomoses64,81,90 are known to influence the thickness of pannus tissue. Moreover, tissue hyperplasia is also a means of compensating for the lack of compliance and contractility in prosthetic grafts. Therefore, variations in flow8,52 and shear stress52 result in an adaptive reduction or increase in thickness of the subendothelial tissue.

16

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.19. Typically meandered edge of transanastomotically outgrowing endothelium (30 µm ePTFE; chacma baboon; 6 weeks).

Fig. 1.20. Loosely arranged endothelial formation at the edge of transanastomotically outgrowing pannus tissue. All the cells are practically lacking contact with the neighboring endothelial cells and are surrounded by a dense carpet of blood platelets (Dacron velour, chacma baboon, 6 weeks).

Apart from these influences, it appears as if the thickness of the anastomotic pannus is to a certain degree also predetermined by an almost template function of the initial fibrinous inner capsule. Even at an early stage when transmural tissue ingrowth has hardly reached the inner fibrinous capsule, pannus formation replicates the dimensions of the fibrin layer which it replaces. Two weeks after implantation for instance, the surface thrombus on ePTFE grafts measures about one-eighth to one-fourth of that on Dacron grafts.1,2,44,60 Similarly, the pannus replacing this thrombus in the anastomotic region not only measures one-eighth to one-fourth on ePTFE grafts compared with Dacron prostheses,2,35 but actually has the same thickness as the preexisting fibrin layer.2,8,64 The alacrity, however, with which this pannus layer grows towards the midgraft region may well be accelerated by the presence of tissue underneath the inner fibrin capsule, as is the case with high porosity grafts. Although the fibrin layer on the blood surface of woven and knitted Dacron grafts is of equal thickness2,45,58,60-62,64-66,91

and transmural tissue ingrowth does not proceed into the inner capsule of either at this stage, pannus outgrowth onto the graft surface seems to be much faster in knitted grafts,2,46,60,68,72,77,81 with their almost continual tissue layer underneath the fibrin capsule compared with the scanty transmural tissue ingrowth beneath the fibrin matrix of woven grafts.40 Similarly, a significantly faster transanastomotic tissue outgrowth occurs in high porosity ePTFE grafts if interstitial tissue development is facilitated through external tissue wrapping. This facilitated outgrowth is observed before transmural tissue ingrowth could account for it.28 As far as the cellular and matrix composition of the pannus tissue is concerned, it appears as if the absence of a mature and confluent endothelium at the time of sub-endothelial tissue development may be one reason for its lack of differentiation. As early as 1975, Mansfield and Sauvage92 described the prevention of undifferentiated pannus outgrowth through in vitro endothelialization of grafts in calves.

The Lack of Healing in Conventional Vascular Grafts

Almost 20 years later clinical in vitro endothelialization demonstrated the same phenomenon: In endothelial lined grafts in which the formation of a confluent endothelium precedes that of the subendothelial tissue,5 smooth muscle cells rather than the occasionally observed fibroblasts61 dominate a mature matrix which even contains an internal elastic membrane.93 The use of purified cultured autologous endothelial cells for the in vitro lining process makes it quite unlikely that smooth muscle cell coseeding was the reason for the development of such a well differentiated subendothelial tissue rather than the interaction with the preexisting endothelium. Since a tissue which lies beneath an internal elastic membrane qualifies per definition as neo-media rather than hyperplastic neo-intima,93 such a differentiation means a huge step towards true healing and therefore needs to be understood with regard to the underlying biological mechanisms. In another study, similar degrees of pannus maturation were observed following endothelial cell seeding with mixed microvascular cells.61 In this case, fibroblasts were initially coseeded with other cell types at similar proportions to those normally seen in anastomotic pannus tissue. Nevertheless, the subendothelial matrix—which matured after the endothelium reached confluence—consisted of mature muscle cells and also elastin rather than the typical poorly differentiated collagen-rich pannus tissue.61 These observations support the presumption that the early presence of an endothelium may influence the degree of differentiation of the subendothelial connective tissue of both pannus and neo-intimal tissue.

Transanastomotic Endothelialization Endothelial outgrowth from the adjacent artery onto the surface of a prosthetic vascular graft is another enigma of prosthetic graft research. For unknown reasons, transanastomotic endothelialization stops shortly after the anastomosis in humans.11,47,60 Yet, the experimental set-up

17

aiming at overcoming this limitation uses animal models which characteristically show premature and rapid surface coverage with anastomotic endothelium. In the past this rapid transanastomotic endothelialization was usually the key parameter of investigations when primarily surface endothelialization was compared between different prosthetic grafts. However, if one agrees that future grafts for clinical use will require transmural tissue ingrowth as a main cell source for surface endothelialization because of their extraordinary length, a better understanding of transanastomotic endothelialization is indeed necessary, but for a different reason than before: To be able to clearly define for each animal model a biological phenomenon which would otherwise continue to blur the distinction line between transmural and transanastomotic tissue ingrowth. Differences Between Animal Models Alerted by the human situation of complete ingrowth stoppage, we can hardly assume that transanastomotic endothelialization is a linearly progressing event in animal models. Therefore, extrapolation of pointwise observations to a long term situation, or even to a constant daily outgrowth distance,2,94 seems obsolete, although it would have provided valuable additional data for an ingrowth curve. Even in the yellow baboon with its vigorous angiogenic capacity, a similar trend of ingrowth stoppage as in humans can be observed: On 30 µm ePTFE prostheses—which do not allow transmural endothelialization—transanastomotic endothelial ingrowth progressively slows down, to stop approximately 2 cm from the anastomosis.95-97 Furthermore, high porosity ePTFE grafts, as well as a few knitted Dacron grafts, cannot be admitted into this comparison because of complete transmural surface endothelialization. Considering these exemptions and the fact that only a minority of studies describe ingrowth margins rather than percentages of surface coverage, it is difficult to compare accurately the various animal

Fig. 1.21. Transanastomotic endothelial outgrowth on 30 µm expanded polytetrafluoroethylene (ePTFE) grafts in four experimental animal models. The curve was plotted using pannus outgrowth values in the dog model (top curve). Points read off the main curve correspond to humans, Papio ursinus/ chacma baboons, and Papio cynocephalus/yellow baboons, as indicated by connecting arrows. The curve was drawn using Deltagraph (for Macintosh).

18

models with regard to transanastomotic endothelial cell outgrowth. Nevertheless, since most of the experiments in the past were based on the canine model, a relatively complete chronology can be estimated for standard grafts in the dog. Single data obtained in other animal models can then be related to the time course of surface healing in dogs. However, the accuracy of even this narrowed appraisal decreases rapidly after the 12th week of implantation when grafts in most of the studies were fully endothelialized and thus can not be used for the assessment of further transanastomotic endothelial cell outgrowth. During the first 3 months of implantation transanastomotic endothelialization in dogs consecutively covers a distance of approximately 2 cm in ePTFE grafts1-3,10,21,86,98-102 and more than 3 cm in Dacron prostheses.2,60,62-64,66,72,76 These ingrowth distances also apply to the second most commonly used animal model, the yellow baboon (Papio cynocephalus).9,19,32,35,52,55,91 Taking into account that the ingrowth margin of endothelium is often very irregular with endothelial tongues reaching far into the graft, a minimal length of 15 cm in all animal models therefore seems a reasonable choice for future 3 month studies. In spite of this general guideline, there is still a distinct difference between the respective animal models with regard to transanastomotic ingrowth behavior. However, since the in vitro proliferative capacity of endothelial cells from humans and experimental animals hardly differs,103 the distinctly prostrate transanastomotic outgrowth potential of human endothelial cells must be a consequence of the complex overall situation of the human model rather than a principal mitotic sluggishness of human endothelial cells. Therefore, one should not downplay the importance of differences between animal models. If one animal model comes significantly closer to the human situation than another, it may well hold the key to understanding the biological reasons for the human failure of graft healing. If transanastomotic ingrowth values of different species on standard ePTFE grafts are related to those of the dog, the coverage which is reached in humans after 56 weeks occurs in dogs after 3.5 weeks,1,2,21,98,86,100 in the yellow baboon after 5.6 weeks and in the chacma baboon after 7.6 weeks. Thus, the transanastomotic outgrowth is 2.1 times slower in the chacma baboon than in the dog. The difference between the two primate species may well be due to the fact that the chacma baboon represents a senescent model with a body mass of up to 40 kg, whereas the yellow baboon is a juvenile model which is on average 2 years old and weighs 10 kg. In this context, it is interesting to note that the difference between the two kinds of baboons is much more pronounced in transmural than transanastomotic ingrowth. Transmural capillarization leads to surface endothelialization of high porosity ePTFE grafts after 2 weeks in the yellow baboon, whereas perigraft tissue has usually not even grown into the outer two-thirds of these prostheses in the chacma baboon (personal observation). Other animal models like rats, calves104 and pigs105 are either too different or not sufficiently documented. However, pigs are usually large enough for vascular implants at a young age, but dog puppies are too small. Therefore, dogs

Tissue Engineering of Prosthetic Vascular Grafts

are usually also senescent at implantation,74 whereas pigs and calves can be expected to be juvenile. In the rat model, the issue of dimensions complicates interpretation. A transanastomotic endothelial outgrowth of approximately 2.2 mm after 6 weeks13,16,17,88 in 30 µm ePTFE makes it impossible to relate these data to the 7.8-13.6 mm of other large animal models. In summary, transanastomotic endothelialization is at least 7.5 times more pronounced in any animal model than in man. Within the experimental set-up, however, the senescent primate model of the chacma baboon at least promises to eventually reveal certain biological principles which make the human healing pattern of prosthetic grafts so unique. Differences Between Grafts Since the endothelial margin represents the frontier of transanastomotically ingrowing tissue,19 it may well be a key player in a synergistic interaction with sub-fibrinous tissue. If one compares the two low porosity prostheses (30 µm ePTFE and woven Dacron), endothelial outgrowth seems to be twice as high on ePTFE1-3,10,21,78,86,98-102 than on Dacron.40,43 This difference is not too surprising, because the mighty inner fibrinous capsule of Dacron seems to have features of an ingrowth barrier for transmural capillaries.43,44,48,59,60,71,72 One can imagine that the same biological reasons which inhibit capillary ingrowth from outside may inhibit endothelial outgrowth from the anastomosis. However, although this fibrin matrix seems to be similar in high porosity2,35,44,60-66,91 and low porosity grafts,45,58 it allows better transanastomotic endothelial ingrowth under high porosity 40,46,47,60,68,72,81 than low porosity conditions.1,2,21,40,43,86,98-102 Since this observation only relates to grafts in which the transmural tissue ingrowth did not traverse the fibrinous capsule and the time would have been too short for fallout endothelialization, the endothelium can be assumed to be of transanastomotic origin. If however, a main difference for transanastomotic endothelial ingrowth on an otherwise not ideal surface seems to be a well developed tissue layer underneath the fibrin capsule, it is reasonable to speculate that this distant tissue layer may somehow affect surface endothelialization.

Porosity Porosity was long known to be a determining factor behind the fate of prosthetic vascular grafts. Initial attempts to replace arteries with solid tubes of synthetic material soon demonstrated that porosity is a prerequisite for graft patency.45,49 Subsequently, pioneers of prosthetic vascular grafts60,71,74 unsuccessfully tried to achieve complete graft healing through porosity. However, only with the advent of highly porous ePTFE grafts did it become apparent that complete surface healing of arterial prostheses was theoretically possible beyond a certain porosity, and this only in a very juvenile animal model. In the endothelium-centered 1980s, excitement arose primarily from the fact that higher porosity may allow surface endothelialization.30,32 However, if tissue engineering of prosthetic vascular grafts ever intends to truly emulate functional arteries, the tandem of deep-seated

The Lack of Healing in Conventional Vascular Grafts

suspicion of the presence of smooth muscle cells and the narrow focus on endothelial cells will need to give way to a more comprehensive mode of tissue ingrowth. Nevertheless, our current biological understanding validates the view that the ability of a prosthetic graft to allow capillary ingrowth still holds the key to today’s grander picture. Researchers were long puzzled as to why interstitial graft spaces of particular prostheses were sufficient to allow macrophages and other inflammatory cells to immigrate, whereas connective tissue cells of similar dimensions were often conspicuously absent.2,18,22,24 Today we understand that certain of these connective tissue cells, like smooth muscle cells, mainly follow the ingrowing endothelial cells.106,107 Therefore, one important determining dimension for complete graft healing is that which mechanically allows capillaries to sprout into the meshwork of a synthetic graft and still leave some space for accompanying cells. Since the average diameter of a capillary is ~10 µm,108-110 it would seem as if the minimum area required for ingrowth of capillaries is at least 20-80 µm.2 The diameter of a functional vessel, i.e., an endothelium and at least one layer of smooth muscle cells, is approximately 23.06 ± 13.11 µm in diameter (personal observations). The attempt to determine porosity requirements for graft healing is no doubt complicated by material specific characteristics and structural uniqueness. The complex three dimensional structure of contemporary grafts results in a wide range of voids between the synthetic surfaces, which hardly allows a precise definition of ingrowth spaces. For manufacturers, a way out of this dilemma was the introduction of water permeability as a means of defining porosity. With this method, materials could be separated on the basis of low and high porosity. For example, materials such as Dacron were distinguished not only on the basis of being knitted and woven, but further characterized as being of high porosity (1500 to 4000 ml/cm 2 /min) or low porosity (200-1000 ml/cm2/min).111 The same principle applied to ePTFE on the basis of distances between the nodes of the expanded material. However, although these values served as useful parameters in characterizing the materials, they were not helpful with regard to healing. The poorly defined space dimensions made it impossible to come to a conclusive answer concerning the role porosity plays in the mitigation of tissue ingrowth into prosthetic vascular grafts. If one attempts to gain a better understanding of the relationship of porosity and healing, one needs to make a clear distinction between porosity and permeability in order to avoid a confusion of terminology. It is important to understand that these terms are not synonymous, because a highly porous material may have a low permeability and vice versa. Permeability The permeability of a material can be defined as its ability to allow other substances (either gaseous, liquid or solid) to pass through its pores or interstices. The customary method for measuring permeability is using water and measuring the area and the applied hydrostatic pressure gradient across a porous membrane. The permeability value is therefore directly proportional to the measured volume of

19

water that passes through the material per unit time, assuming that the thickness of the membrane remains constant. The resultant measurement has been misleadingly called “water porosity” by surgeons and manufacturers alike. Porosity Porosity, in contrast, refers to the void spaces (pores) within the boundaries of a solid material, compared to its total volume. Although one could argue that the voids within a material will allow water to flow through them, numerous pores may well be cul-de-sacs and thus fail to provide an open corridor from one side of the graft to the other. Moreover, vascular prostheses are not designed to be rigid structures. They are compressible for ease of handling and can be readily deformed by radial and longitudinal tensions as well as by internal pressures, even in the relatively noncompliant contemporary grafts. Such stresses not only change the dimensions of the prostheses, but also lead to a redistribution of the yarns and fibers within the material. Naturally, the change in textile structure by both circulatory stress and compression by ingrowing tissue itself affects the distribution, size and tortuosity of the channels that run from one side of the graft to the other. Since the limiting factor for tissue ingrowth is the bottleneck of the narrowest part of a transmural space, it is further important to look at interstitial spaces throughout their three dimensional course. Unfortunately, hardly any information is available regarding the dimensions of porous structures in vascular grafts. In order to get a better insight into this critical parameter for tissue incorporation, we have evaluated interstitial space dimensions in a few typical contemporary grafts, both before and after implantation. Moreover, we have investigated these prostheses regarding their true maximum interstitial ingrowth spaces and their resulting theoretical ability to accommodate transmural microvessels. Microvessels were decided to be the dimension against which measurements were compared, based on their central role in a variety of healing aspects from surface endothelialization to neo-media formation. Theoretical Considerations for Healing of Vascular Grafts Although the process of expanding teflon material to the typical nodes-and-fiber structure of ePTFE grafts (Fig. 1.22A) allows for a wide range of internodal distances, the resulting changes in ingrowth spaces are moderate. With well-defined anchor points and a rigid material like teflon, the distensibility of the fibrils is limited (Fig. 1.22B), thus making disproportional increases in internodal distance necessary to achieve increases in fibril length and hence an increase of ingrowth dimensions of a few micrometers. To try to correlate true ingrowth spaces with internodal distances as defined by the manufacturers of ePTFE grafts is insufficient, therefore, if not combined with the assessment of fibril dimensions. Furthermore, in order to accurately determine ingrowth spaces, one needs to assess these structures throughout the entire thickness of the graft wall. In a similar way, one needs to measure more than just bundle densities and fiber diameters in knitted (Fig. 1.24) and woven

20

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.22. Scanning electron micrograph of a 30 µ m expanded polytetrafluoroethylene (ePTFE) graft (Impra Inc., USA) showing the typical nodes-and-fibril structure. x750, (top) inner surface and (bottom) outer surface.

Dacron materials (Fig. 1.25) to obtain an accurate measurement of cell ingrowth spaces. These assessments are further complicated by the introduction of externally wrapped PTFE grafts (Fig. 1.23) and veloured Dacron surfaces (Fig. 1.26).60 These coverings are intended to prevent the grafts from aneurysmal dilatations and to increase the tissue bond between graft and perigraft tissues, with a subsequent better healing response.111 Taking all these factors into account, our approach toward the assessment of true ingrowth spaces in commercially obtained both low porosity (< 30 µm) and high porosity (60 µm) ePTFE, Dacron velour and knitted Dacron

grafts was based on the combined image analyses of histological cross-sections and scanning electron microscopy (SEM) of surface structures (see figure legends). Based on these measurements (Table 1.1), we attempted to reconstruct the maximum ingrowth space dimensions throughout the graft wall and draw conclusions about whether the graft materials would sufficiently allow transmural ingrowth for microvessels invading from the adventitia. In order to calculate ingrowth spaces, we measured a number of parameters related to the material characteristics (Table 1.1). Scanning electron microscopic examination of the two different porosities of ePTFE used in this study gave a good

The Lack of Healing in Conventional Vascular Grafts

21

Fig. 1.23. Scanning electron micrograph of a 30 µ m expanded polytetrafluoroethylene (ePTFE) graft (Impra Inc., USA) showing external circumferential wrap. The nodes-andfibril structures can be seen through the wrap. x750, outer surface.

Fig. 1.24. Scanning electron micrograph of a knitted Dacron prosthesis (Vascutek; water porosity = 1350 ml/ cm2/min), showing reasonably tightly knitted bundles of fibers, x75.

indication of surface dimensions with respect to internodal distances, internodal spaces, mean fibril lengths and mean interfibril space (Table 1.1). Internodal distance was measured from the center of one node to the center of another, whereas internodal space width was defined as the actual space between the margins of two adjacent nodes. No significant difference was found between the inside and outside surfaces for each respective graft. However, the internodal distance stated by the manufacturers did not precisely match the internodal distance we measured. It could be argued that the internal diameter of ePTFE grafts may influence the manufacturing process and thus

the node-to-fibril structure and dimensions. Based on this, we measured internodal fibril lengths in 6 contemporary 30 µm ePTFE grafts with varying internal diameters (Table 1.2). Only the inner surface was measured, as all the grafts contained dense outer circumferential wraps sometimes additionally reinforced with rings. The results showed a reasonably consistent fibril length irrespective of internal diameter (Table 1.2). The values of 24.28 ± 3.48 µm and 20.69 ± 2.35 µm for the 6 and 10 mm I.D. grafts (W.L. Gore and Associates, Flagstaff, Ariz.) respectively, are in agreement with earlier studies.1 However, although there were hardly any differences in internodal distance between grafts of

22

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.25. Scanning electron micrograph of a woven Dacron prosthesis (DeBakey; water porosity = 232 ml/ cm 2 /min), showing tightly woven, well-defined bundles of fibers, x75.

Fig. 1.26. Scanning electron micrograph of the double veloured Dacron prosthesis (Cooley; water porosity = 1660 ml/cm2/min), showing a completely wide open, random array of fibers. Inset: Dacron felt (USCI) used in our study, x75.

different inner diameter from the same manufacturer, there were distinct differences in internodal distance between comparable grafts of different manufacturers. In 30 µm ePTFE grafts from Impra (Table 1.1), a mean inner fibril length of 17.79 ± 5.62 µm was opposed by 23.57 ± 6.21 µm measured in W.L. Gore products (Table 1.2). In order to calculate the minimum continual transmural ingrowth channels through the wall of both ePTFE grafts (Figs. 1.27 and 1.28), a mathematical approach was chosen relating three dimensional structures of nodes and fibers to factors such as the obstruction caused by internodal fibrils from one layer to the next. The underlying as-

sumption of a worst scenario was that, beginning at the adventitial surface, each successive layer of fibrils lies in the center of the preceding layer, thereby creating the greatest possible obstruction. Therefore, this approach provides the narrowest possible channel spaces for continual tissue ingrowth from outside to inside. Figures 1.27 and 1.28 illustrate the theoretical comparison between the minimum continual transmural ingrowth channel and the maximum available ingrowth space of a given layer throughout the walls of the 30 and 60 µm PTFE grafts. The reason for the discrepancy between the two curves in the 30 µm PTFE (Fig. 1.27) is the fact that the

The Lack of Healing in Conventional Vascular Grafts

23

Table 1.1. Material characteristics of 5 mm internal diameter ePTFE grafts (Impra, Inc.) Internodal Distance as stated by the manufacturer (mm)

Mean Internodal Distance (peak to

Mean Internodal Space Width (mm)

Mean Fibril Lengths (mm)

Mean Interfibril Space (mm)

30:INSIDE

24.12 ± 6.47 27.27 ± 13.19

60:INSIDE

53.01 ± 20.44

60:OUTSIDE

47.94 ± 9.90

17.79 ± 5.62 95% CI: 14.92-18.84 15.14 ± 6.12 95% CI: 12.95-17.33 45.72 ± 16.22 95% CI: 40.27-51.17 30.80 ± 7.03 95% CI: 28.43-33.16

ND*

30:OUTSIDE

16.88 ± 5.49 95% CI: 15.78-19.80 14.49 ± 6.21 95% CI: 12.27-16.72 42.41 ± 15.39 95% CI: 37.23-47.58 28.14 ± 6.36 95% CI: 25.99-30.28

5.11 ± 2.39 ND* 4.53 ± 0.76

Material characteristics of both low (30 µm) and high (60 µm) porosity 5 µm internal diameter (ID) ePTFE grafts (Impra, Inc.) as measured by SEM anaylses. Counts were done over 5 random fields using a calibrated NIH image analyses system. Results are presented as mean ±SD with 95% confidence intervals calculated around each mean vaule. n = 60. ND* = Not determined beacause cell ingrowth occurs through the fibrils from the outside surface only.

Table 1.2. Comparative values of contemporary 30µm ePTFE grafts (W.L. Gore and Associates) 30mm ePTFE Internal Diameter (mm)

Mean Internodal Fibril Length ± SD (mm) (Inner Surface)

3 5 6 10 15 22

21.12±5.56 23.57±6.21 24.28±3.48 20.69±2.35 32.09±5.10 26.17±2.43

Counts were done on 5 random fields on the inner surface of all the grafts using calibrated NIH image analysis software. The outer surface of all the grafts was covered with a circumferentially dense wrap. The 5 and 15mm ID grafts were further reinforced with external rings. Results are presented as mean±SD. A clear discrepancy exists between the 5mm ID internodal fibril lengths of the two manufactured grafts.

Fig. 1.27. Comparison of mean ingrowth space and the theoretically predicted minimum continual transmural ingrowth channels for 30 µ m ePTFE. The ingrowth spaces (■) were determined by image analysis on histological cross-sections (n = 2) which were subdivided into 10 consecutive layers extending from the lumen to the adventitia. In each layer, internodal measurements were done in 8 random fields using a Leica Q500MC image analyzing system. Mean ingrowth spaces are expressed as graphic points for each measured layer (µm). Vertical lines denote ±1 S.D. For the theoretically calculated continual transmural ingrowth channels (–◊–), internodal spaces, fibril lengths and mean interfibril spaces were measured using scanning electron microscopy. The results were plotted on a logarithmic scale representing full wall thickness (2000 layers). Vertical lines denote ±1 S.D.

24

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.28. Comparison of mean ingrowth space and the theoretical prediction of the minimum continual transmural ingrowth channel for 60 µm ePTFE. The identical protocol was followed as in Figure 1.27.

maximum space (Fig. 1.27, line marked with solid squares) has been worked out on the basis of the maximum ingrowth space available between two fibrils without taking the interaction of the neighboring fibrils into account. These spaces are therefore ingrowth spaces in isolation. If one compares the maximum ingrowth space available through the graft (assuming that each histological layer equals 200 true layers as previously established by SEM, resulting in a total of 2000 layers throughout the entire wall thickness) to the theoretically available transmural ingrowth channel (Fig. 1.27, lined marked by open diamonds), it becomes obvious that the available space is in fact much less than it appears initially on the basis of isolated layers of nodes and fibrils. Interestingly, the minimal ingrowth channel from outside for 30 µm PTFE grafts progressively diminishes in size over the first 10 layers, as a result of the obstruction caused by the successive layers of fibrils. Thereafter, the values remain consistent before increasing again till layer 200. Throughout the rest of the graft, the predicted channel size would not accommodate capillaries, although the ingrowth space of each isolated layer of nodes and fibrils would remain theoretically sufficient for capillary ingrowth. This calculation corresponds with the histological observation that only a few capillaries are found in the wall of nonwrapped 30 µm PTFE grafts, and those few are restricted to the outer third. In 60 µm PTFE grafts (Fig. 1.28), the minimum dimension of transmural ingrowth channels from outside also progressively diminishes in size in the first 20 layers, after which the values become consistent throughout the rest of the graft wall till layer 200. From then onwards the maximum diameter of a continual transmural ingrowth channel decreases again as a consequence of the narrowing of the internodal space. Despite the fact that the calculated channel size represents the “worst” scenario in terms of ingrowth dimensions, as it nears the lumen it still provides sufficient space for ingrowing arterioles within the 5% confidence interval. Therefore, complete transmural vascularization would still be possible in these grafts. Transmural capillarization,

in contrast, would be permitted by 100% of the transmural spaces. Assessment of Ingrowth Spaces in Dacron Grafts If the task of assessing interstitial graft spaces is considered difficult in ePTFE, it gets even more challenging when Dacron prostheses are concerned. In the era prior to protein pretreatment of vascular prostheses, porosity was directly proportional to bleeding. Therefore, textured grafts made of polyester braids aimed at providing surgeons with an instantly hemostatic alternative. In contrast to ePTFE, however, high or low porosity is not primarily achieved through a variation of the same manufacturing process, but rather through the different procedures of knitting and weaving. Moreover, an additional variable is introduced by the needle density used in the knitting process and by optional combination with a velour surface, which further influences the size of the open spaces between the braids. Woven Dacron Commercially, woven fabrics consist of two sets of yarns, the warp and the weft, which are interlaced at right angles to each other (Fig. 1.25). Woven material is often preferred by surgeons, as it is known for its high bursting strengths, low permeability to liquids, minimal tendency to deform under stress and less proneness to kinking. However, the woven structure also has poor compliance, limited elongation, a tendency to fray and, above all, a very low porosity. In spite of this low porosity, tissue ingrowth into the tight interstitial spaces of the fabric is occasionally observed.40,57 However, since the available space does not appear to allow for cell infiltration in the manufactured configuration, one must assume that cellular ingrowth is a consequence of a rearrangement of the relatively loosely lying interstitial fibers. Theoretically, therefore, the number of fibers, fiber diameter and bundle area could be measured to obtain a reasonable measurement of ingrowth spaces. This information

The Lack of Healing in Conventional Vascular Grafts

would then be used to determine a packing factor for the material. The main shortfall of this approach, however, is that although the native graft material seems flexible and reorganization of fibers seems theoretically feasible, the interstices of the graft upon implantation are soon filled with a fibrin matrix which polymerizes into a fairly rigid structure, thus embedding the fibers and bundles and allowing very little room for movement. It seems likely that the proteolytic process required for this rearrangement needs a critical tissue presence with its associated degrading capacity, which in turn would need a higher initial porosity. This may explain why single cells such as macrophages and fibroblasts are found near the surface of woven prostheses, but only very occasionally capillaries.40,57 Knitted Dacron According to the manufacturers,112 knitted constructions contain a set of yarns which are interlooped around each other as opposed to interlaced (Fig. 1.24). The specific type and the compliance behavior of knitted fabrics depends on the direction in which the yarns are interlooped. If the yarns lie predominantly in the lengthwise direction then the fabric is “warp knit” whereas if they lie in the transverse direction, the fabric is “weft knit”. A significant difference between these two types of knit is that weft knit fabrics will unravel, whereas warp knits will not. This is an important consideration for the surgeon, as the grafts are often cut at preimplantation for perfect abutting to the anastomoses. Undoubtedly, once implanted, the difference is negligible. Due to the difference in structure of Dacron grafts as opposed to PTFE, a slightly different approach was adopted to calculate ingrowth spaces. Surface measurements made by SEM, and the histological assessment of wall thickness and ingrowth space dimensions, are the same as in ePTFE. However, while fibrils cause the main spatial limitation in PTFE, fiber orientation and the arrangement of fiber bundles throughout the graft thickness limit the spaces in knitted

25

Dacron. We therefore analyzed explanted histological sections for knitted as well as pure velour prostheses (Fig. 1.26) through ten layered areas from the adventitial surface to the lumen. The details of the measurements are described in the figure legends. The commercial knitted graft (Vascutek) used in this study had a water porosity of 1350 ml/cm2/min, falling into the lower range of porosities for knitted materials.15 On the basis of interfiber measurements through the graft wall, a graphical representation showing the mean values for maximum ingrowth spaces could be obtained (Fig. 1.29). In spite of the fact that this material lies in the lower range of porosity within knitted grafts, it was still surprising to see how narrow the maximum ingrowth spaces through the knitted graft were. When one compares the percentage distribution of spaces theoretically permissive for capillaries and/or arterioles (Fig. 1.29B,C), however, the mean size of spaces within the first third of the graft wall (14.44 ± 3.74 µm, layers 7-10) would accommodate capillaries. In addition, albeit at a low percentage (5.16 ± 0.56%), the mean ingrowth space of 24.64 ± 1.75 µm would also allow a few arterioles to surpass this zone (Fig. 1.29C). Towards the middle of the graft (layers 4-6) the restriction for capillaries becomes less pronounced (13.92 ± 2.68 µm), defying the common sense expectation that cells would encounter more tightly packed bundles in this midzone of the prosthesis. The luminal third of the graft (layers 1-3) has consistently narrow spacing, allowing capillary ingrowth only in 30.4 ± 0.38% with a mean ingrowth space of 12.34 ± 2.50 µm. None of the spaces in this zone allow arteriole ingrowth (Fig. 1.29C). Velour Surfaces The term “velour” in a textile context refers to a thickbodied fabric, either woven, knitted or nonwoven, that has a smooth, soft surface by virtue of additional yarns. All commercial velours, however, are knitted designs. They may be distinguished from the earlier knitted prostheses by their Fig. 1.29. Graphic representation of maximum ingrowth spaces for knitted Dacron. Histological cross-sections of explanted samples were subdivided into 10 consecutive layers extending over the full thickness of the wall from lumen to adventitia. Interfiber spaces were measured in 8 randomly selected fields in each zone. Results in (A) are expressed as mean ± SD. A frequency distribution histogram was applied to the results from (A) to obtain percentages of ingrowth spaces of > 10 µm and > 23 µm (B, C).

26

thicker fabric and greater porosity, as well as lower fabric density and packing factor. To assess ingrowth space within velour graft types has been very difficult, and most researchers have relied on methods such as water porosity. However, a few authors113 have attempted to measure its true dimensions by looking at pile height and the amount of piles in the veloured knit. The impetus for using velours stemmed from a conviction that a rougher and more porous surface both externally and internally would facilitate the development of neointimal and perigraft tissue. In order to assess the qualities of pure velour structures, we manufactured velour tubes made of Dacron velour felt (Fig. 1.26, inset). In sharp contrast to the seemingly restrictive interfiber spaces of the knitted material, the Dacron velour showed mean ingrowth spaces ranging from 30.02 ± 2.95 µm at the adventitia, to 36.88 ± 20.06 µm at the lumen, with an overall mean ingrowth space of 27.97 ± 4.80 µm (Fig. 1.30A). These mean ingrowth spaces, as in the case of the knitted grafts, were assessed on the basis of histological measurements of interfiber distances through ten consecutive areas of explanted samples (see figure legend). Interestingly, although the mean space seems sufficient for arteriole ingrowth, the wall profile (Fig. 1.30A) shows that the space becomes narrowed towards the central part of the graft and may have a restrictive effect on ingrowth. The reason for this central restriction is the presence of numerous bundles in this area causing an overall smaller interfiber space. Despite this restriction, the calculated maximum spaces would still allow the ingrowth of capillaries and arterioles at all levels. Percentage distribution profiles for both the capillaries and arterioles (Fig. 1.30B,C) show a similar distribution pattern as the mean ingrowth space profile (Fig. 1.30A), with the majority of spaces of more than 23 µm (Fig. 1.30C) being in the outer and inner one-third of the graft. Taking all these results into account, it seems that the dimensions of graft porosity may have been overestimated Fig. 1.30. Graphic representation of ingrowth spaces for Dacron velour. An identical protocol was followed as for Figure 1.29. Results are expressed as mean ± SD.

Tissue Engineering of Prosthetic Vascular Grafts

for most of the contemporary vascular prostheses. It becomes clear that, although surface measurements of graft materials give a good indication of ingrowth characteristics, additional information such as interfiber distances and theoretically predicted channel spaces throughout the graft wall are critical for assessing the degree of pure mechanical restrictions for tissue ingrowth into prosthetic vascular grafts.

Fibrin Prior to implantation, conventional nonpretreated grafts represent simply synthetic scaffolds. Therefore, all biological components which eventually fill the interstices of these scaffolds are host derived. Within this process of host incorporation, proteinaceous ingrowth matrices naturally precede infiltration with cellular components. Since the implantation of a vascular prostheses inflicts a surgical wound, this proteinaceous matrix is fibrinous. In low porosity grafts, the surrounding fibrin coagulum derives entirely from iatrogenically injured vessels. The cut ends of these vessels constrict within a few minutes and get sealed with aggregating blood platelets. As a result, the hematoma around the graft usually remains moderate. Within a few minutes tissue thromboplastin generates thrombin, which, together with platelet factor 3 (PF 3) released by the few entrapped platelets, leads to the transformation of fibrinogen into fibrin. Particularly in nonpretreated porous grafts, blood may also extravasate from the graft lumen to the outside. In this case the activation of the clotting cascade occurs through contact of factors XII, prekallikrein, XI and high molecular weight kinogen114 and, again, through released platelet PF 3. Although blood does not seep through the wall of less porous grafts, plasma still clots within the synthetic meshwork through a similar mechanism. Therefore, one can assume that the entire fibrinous matrix initially surrounding the graft and filling its interstices is representative of typical wound fibrin, which distinctly favors tissue ingrowth.

The Lack of Healing in Conventional Vascular Grafts

27

Fig. 1.31. Hypothetical sequence of cellular ingrowth events in native fibrin. The normal hemostatic process of clot formation leads to a relatively low platelet number, fibrinogen and TGF-β content. The resulting low density of thick fibrin fibers together with mild hypoxia has a pro-angiogenic effect as it causes an immediate upregulation within the endothelial cell of integrins αvβ3 and αvβ5, which facilitate adhesion to the extracellular matrix, and the release of proliferative and pro-migratory cytokines such as vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF) and basic fibroblast growth factor (bFGF). The upregulation of the endothelial cell-specific integrins causes the release of serine proteases, urokinase and tissue plasminogen activator (u/tPA). These proteases in turn convert surrounding plasminogen to plasmin, which acts on the neighboring extracellular matrix to expose attachment sites for the cell-specific integrins. This whole process eventually leads to revascularization and the maintenance of vessel integrity at the wound site.

In contrast to this loose initial fibrin matrix, an entirely different fibrin clot may be continually generated on the blood surface. While the interstitial and external clotting environment is characterized by a small blood pool with limited availability of coagulation factors and platelets, undiminishing amounts of fibrinogen and platelets are present at the blood-contacting surface of the prosthesis. Therefore, this consecutively increasing inner layer of thrombus obviously differs in both its mode of generation and its composition. Outer Fibrin Matrix Since tissue ingrowth into prosthetic grafts occurs primarily from outside, the outer fibrin clot surrounding the grafts serves as a provisional matrix for cell migration from the adjacent tissue. However, in order to understand possible adverse qualities of the inner fibrin matrix, one needs to understand the principal interactions occurring between cells and the more natural wound-type fibrin on the outside of the graft. This blood clot filling the adventitial wound space consists mainly of crosslinked fibrin, plasma fibronectin, vitronectin and thrombospondin (TSP),115 containing low levels of platelet derived cytokines like TGF-β and PDGF. This fibronectin-rich fibrin matrix offers binding sites for integrins of ingrowing cells. For macrophages it is primarily α5β1,116 whereas it is αvβ3 for new blood vessels.117 The integrin αvβ5 has also been implicated in the formation of new blood vessels.118 However, whether invasion of endothelial cells into a fibrin matrix is αvβ5-dependent remains to be established, as the role of this integrin points to it interacting selectively with a vitronectin-rich matrix.119 The interaction between integrins and the matrix also influences other regulatory mechanisms. For instance, a “stressed” fibrin matrix as a result of contraction increases

the responsiveness of cells to PDGF.120 This is an important aspect of the triggering of angiogenesis at such an early stage. Furthermore, TGF-β—which is a chemotactically highly potent cytokine for macrophages—is potentially active in fibrin,121 whereas it is inactivated in collagen matrices by decorin.120 The relatively low platelet concentration in wound fibrin makes it likely that TGF-β augments macrophage infiltration without inhibiting vessel ingrowth. This can be explained by the concentration dependence of TGF-β, which is only inhibitory for angiogenesis at high concentrations 108 but chemotactic for macrophages in trace amounts.122 This early fibrin matrix also facilitates protease upregulation and the resulting degradation of the fibrin clot to enable tissue ingrowth. The initial events in this process are again regulated through an integrin-ligand mechanism. Ingrowing endothelial cells for instance, induce their protease secretion through the binding of α5β1 to fibronectin and fibronectin fragments in the fibrin matrix.117,120 The relatively low number of platelets contributing to this initial clot also makes it likely that the concentration of plateletderived thrombospondin (TSP1)—a potent inhibitor of proteases145—does not result in a mitigation of tissue ingrowth. Inner Fibrin Matrix Although the mechanism of ingrowth inhibition into prosthetic vascular grafts may well be a multifactorial event, it is conspicuous that the most impenetrable part of the graft wall in humans seems to be the inner fibrinous capsule. This phenomenon has been reported for more than 20 years in porous Dacron grafts.43,44,48,59,60,71,72,74,75 We have also observed it in 60 µm ePTFE prostheses in the chacma baboon where the impenetrability of the inner fibrin layer even applies to inflammatory cells. After a few weeks of implantation, we found the fibrin coverage of the blood surface to be

28

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 1.32. Hypothetical sequence of events in the inner fibrin capsule on Platelet Release prosthetic grafts. Fibrin is continually deReaction posited due to platelet and macrophage accumulation and their subsequent degranulation. Platelets and red blood cells become entrapped in a constantly increasing, compacted fibrin layer. Degranulation of the α-granules of the Inactive TGFβ platelets causes a microenvironmental Mo Tissue Factor active TGFβ increase of inactive transformting growth factor-beta which becomes sePLASMIN questered into the extracellular matrix in PLASMINOGEN a latent form. Endothelial cells at the outer surface of the graft material Migration UPA/tPA α β α β upregulate u/tPA through an integrinProliferation v 3 5 1 extracellular matrix mediated effect. Differentiation These serine proteases convert bloodborne plasminogen to plasmin, which in turn activates the latent TGF-β. TGF-β acts on the endothelial cells to initiate differentiation with a concomitant decrease in migration and proliferation. This mechanism may therefore cause a premature cessation of angiogenesis and a reduction in transmural vascularization. Additionally, blood-borne macrophages express tissue factor following adherence to biomaterials. This transmembrane protein initiates further fibrin deposition. Macrophages in the depth may shed vesicles containing tissue factor activity which could further facilitate fibrin deposition on the surface.

almost completely acellular (Fig. 1.10). It is therefore tempting to ask whether it is possible to recognize obvious differences in the composition or structure of the inner fibrin capsule which may be responsible for this distinct ingrowth inhibition. In pursuing this question it seems as if the mode of fibrin generation on the inside of grafts already provides significant clues with regard to growth inhibitory properties. First of all, there is a uniqueness about this particular fibrin clot: Both clotting factors and platelets are not consumed from a predefined quantity as is the case in a hematoma, but rather replenished on an ongoing basis. On the one hand, a process of platelet-enrichment through surface adherence results in an disproportionate density of thrombocytes which undergo activation through surface contact and shear stress. As a consequence of this augmented platelet release reaction, high concentrations of α-granule products like platelet factor 3, fibrinogen and von Willebrand factor, as well as dense granule contents such as calcium, continually sustain the coagulation process.123-125 On the other hand, the rapid replenishment of blood plasma next to the platelet carpet on the graft surface further guarantees undiminishing concentrations of fibrinogen and other clotting factors. Thus, one can assume that the fibrin generated on the inner surface of vascular prostheses has a high content of fibrinogen. Since TGF-β is also liberated from α-granules, and its incorporation into fibrin matrices correlates with its preexisting concentration, this fibrinogen rich fibrin matrix can also be expected to be particularly rich in TGF-β. Based on these considerations, one needs to ask the initial question differently: Can high fibrinogen concentrations as well as other α-granule release products like TGF-β and thrombospondin and dense granule products like calcium

affect the structure and/or composition of fibrin in such a way that it becomes hostile to tissue ingrowth? Fibrin Structure It is a relatively recent discovery that fibrin structure influences cell growth. At lower fiber density a fibrin matrix strongly stimulates capillary morphogenesis,126 whereas it only stimulates endothelial migration without three dimensional capillary formation at higher fiber density. On an ultrastructural level, a similar inverse correlation between cell growth and fiber density was found.127 The higher the density of fibrin fibers and the lower the average thickness of these fiber bundles, the lower the penetrability for cells.127 Since in vascular grafts cellular ingrowth seems more inhibited in the inner than the outer fibrin matrix, it is therefore of interest whether inside circumstances favor the generation of a denser fibrin matrix with lower fiber bundle thickness. It was recently demonstrated that an increase in fibrinogen concentration by a factor of 3 resulted in a decrease in fiber bundle thickness and an increase in fiber bundle density, also by a factor of 3.127 Considering the high platelet density on the blood surface of grafts and the undiminishing pool of clotting factors, one can assume that the fibrinogen concentration inside this surface clot is higher than outside, thus resulting in a fibrin matrix of higher fiber density and lower fiber diameter. The suspicion that higher fibrinogen contents inhibit cellular ingrowth was further substantiated by other studies using fibroblasts128 and endothelial cells.61,129 Cytokine Content The high density of platelets adherent to the blood surface, as well as their ongoing activation,5,130 suggest a high

The Lack of Healing in Conventional Vascular Grafts

concentration of platelet α-granule contents at the time of coagulation. Amongst other substances, the α-granules of platelets contain platelet derived growth factor (PDGF) and transformting growth factor-beta (TGF-β).5 PDGF is a dimeric growth factor which is mitogenic for endothelial cells and smooth muscle cells. TGF-β, in contrast, is a potent inhibitor of cell proliferation, migration and proteinase production in vascular endothelial cells.131 In view of the suspected inhibitory effect of the inner fibrin matrix on angiogenesis it is reasonable to hypothesize that particularly high levels of TGF-β contribute to the mitigated vessel outgrowth. However, the TGF-β content of human platelet αgranules exists exclusively in the β1 isoform, stored in a latent form in which the precursor peptide (the latency-associated peptide LAP) remains noncovalently associated with the mature 25 kDa TGF-β1 dimer.132 In addition, the latent TGF-β complex from platelets has an additional protein, the latent TGF-β binding protein (LTB-P) which is linked with a disulphide bond to the LAP.133 Since the activation of TGF-β only occurs when the mature 25 kDa dimer is released from the LAP121 and the entire complex is incorporated into the fibrin structure, fibrin degradation is a prerequisite for liberating inhibitory TGF-β concentrations from the matrix. Coincidentally, the release of the active 25 kDa TGF-β dimer needs the same proteases, such as plasmin, which are needed to degrade the fibrin matrix.134 Therefore, a likely explanation for the late inhibition of capillarization through TGF-β would be the following scenario: Initially, endothelial cells from the adventitial region of the vascular prosthesis are attracted into its interstices by the PDGF and VEGF released by platelets and polymorphnuclear granulocytes.135,136 Since the fibrin matrix represents a more or less normal wound matrix at this stage, cell receptors involved in initial migration and proliferation, such as the integrins αvβ3 and αvβ5, are expressed together with the upregulation of serine proteases (uPA, tPA), enabling the matrix degradation required for migration. Moreover, the interstices of the prostheses contain a mixture of macrophages and foreign body giant cells (FBGC) which are generators of large amounts of pro-angiogenic cytokines such as bFGF (predominantly under hypoxic conditions), and only minor amounts of inhibitory TGF-β.56 Under those favorable circumstances, capillary ingrowth slowly proceeds towards the inner fibrin layer with its higher fibrinogen and TGF-β content, as well as its denser fibrin structure. With increasing TGF-β concentrations as well as changing fibrin composition, angiogenesis slows down. The increase of TGF-β may be further augmented by macrophages, which degrade fibrin matrices through the upregulation of endogenous serine proteases without concomitant release of plasminogen activator inhibitor (PAI-1). Therefore, matrix degradation by macrophages is much less controlled than degradation by endothelial cells, where a balanced, focalized proteolysis of the pericellular matrix through the up or downregulation of proteases and PAI-1 occurs.137,138 Eventually, both fibrin structure and liberated and activated TGF-β amounts may be sufficient to bring ingrowing capillaries to a complete halt. A main challenge in substantiating this scenario will be the detection of high TGF-β amounts in the inner fibrous capsule, because the

29

incorporation into the LAP-LTBP-TGF-β complex may well make it immunohistochemically unrecognizable. Inhibition of Matrix Degradation Apart from directly inhibiting ingrowing endothelial cells, capillarization and tissue ingrowth into the inner fibrin layer may also be indirectly inhibited through a particularly high PAI-1 concentration in the matrix. Fibrinolysis of the fibrin matrix in normal wound healing is primarily due to plasmin activation from its precursor molecule, plasminogen. In vivo, this conversion to plasmin is mediated by tissue plasminogen activator (tPA) and urokinase plasminogen activator (uPA), both of which are derived from endothelial cells, macrophages and giant cells. The action of these proteases is inhibited by PAI-1, secreted by endothelial cells and platelets. If one assumes that platelets play an disproportionate role in fibrin generation on the inside of the graft, one must also reckon with an disproportionately high level of platelet-derived PAI-1.139 Angiogenesis Inhibitors Another platelet-derived protein which may detrimentally affect capillarization of the inner fibrin capsule is the extracellular matrix molecule thrombospondin (TSP1). TSP1 is a large, multi-functional ECM glycoprotein that can influence endothelial cell function in vitro140 and angiogenesis both in vitro and in vivo.141,142 The role of TSP1 in modulating endothelial cell function has been widely investigated. TSP1 was initially identified and characterized as a protein released from the α-granules of platelets upon activation by thrombin. TSP1 is now known to also be produced by many other cell types such as endothelial cells, fibroblasts, smooth muscle cells and monocytes/macrophages. Thrombospondin occurs in two forms—soluble and matrix-bound. The effect of the matrix-bound TSP1 is often different from that of the soluble form. Soluble TSP1, for instance, has been demonstrated to inhibit the proliferation of endothelial cells143 while matrix-bound TSP1 appears to be a permissive substrate for endothelial cell proliferation. Migratory, invading endothelial cells adhere to insoluble, matrix-bound TSP1 while soluble TSP1 may saturate surface receptors and produce an anti-adhesive effect.144 The anti-adhesive activity of soluble TSP1 may interfere with the angiogenic process by preventing the cell-to-substrate or cell-to-cell interactions necessary for endothelial cell migration and capillary formation. Furthermore, recent reports have revealed that TSP1 can also function as a protease inhibitor.145 These findings suggest that, apart from its direct effect on cells, soluble TSP1 may also influence angiogenesis by affecting ECM turnover and composition. There is further evidence that TSP1 may inhibit the proteolytic enzymes of the fibrinolytic pathway, including plasmin and urokinase plasminogen activator.146 The modulation of protease activity therefore provides a potentially crucial role for TSP1 in regulating angiogenesis, as it seems to be a factor in correcting the balance of protease activity, which is essential for angiogenesis. Another important interaction of TSP1, particularly with respect to capillarization of vascular prostheses, is its

30

interaction with transformting growth factor-beta (TGF-β). TSP1 has been shown to bind and activate secreted and ECM-sequested transformting growth factor-b.147 It follows therefore that the interaction between soluble TSP1 and TGF-β may prove to be crucial in inhibiting cell growth and proliferation. Even more interesting is the fact that TGF-β and TSP1 are both found in the α-granules of platelets,148,149 increasing their presence in equal proportions in an environment characterized by a significant platelet release reaction. Since thrombospondin released by platelets is not incorporated into a secreted extracellular matrix, it would provide another explanation for the inhibitory effect of a platelet-rich fibrin matrix, since the anti-proliferative effect of TSP1 is determined by the soluble form of this secreted protein. The one paradox that remains unresolved is the finding that macrophages in wounds, as well as in other inflammatory settings, actively produce TSP1. The finding that the macrophages are releasing a molecule that inhibits angiogenesis is at first surprising. However, TSP1 release from macrophages may represent a “direct” control of angiogenesis, instead of an “indirect” control of angiogenesis through the protease activation of latent TGF-β. Fibrinogenic Effect of Macrophages One of the most puzzling questions is why the inner fibrin capsule of Dacron grafts differs from that of ePTFE grafts although the blood exposure of the material itself is only a transient initial event. From the first thin layer of fibrin which covers the synthetic material, one would expect no further difference between the surfaces with regards to the thrombus formation. This, however, is not the case. If the exposed surfaces themselves do not differ because of the capping effect of the fibrin, what are the most likely differences which could be responsible for the qualitative and quantitative composition of the inner fibrin capsule? One such difference may well be the surface structure. Prior to any cell accumulation, the relatively coarse surface of knitted Dacron is covered by a thrombus of more than 100 µm,44,60 whereas the fine-structured surface of ePTFE lies under a thin fibrin coagulum of only 15 µm.1,2 However, the fate of crimped surfaces demonstrates that fibrin tends to equalize surface unevenness. As a consequence, all graft surfaces will eventually have the same smooth fibrin coverage, irrespective of their structure. If there are no differences between the initial surface fibrin, and the material itself is hidden underneath, the most likely explanation for an ongoing difference in thrombogenicity would be the active participation of cellular components in the coagulation process. Since the predominant cell type at the time of the most significant difference in fibrin buildup between the Dacron and ePTFE are inflammatory cells, and later on specifically the macrophages, one needs to answer the question of whether macrophages may be key players in the formation of the inner fibrin capsule. In this context, it is interesting to note that macrophages are able to express procoagulant activities (PCA) through their adherence to surfaces.150 Adhesion of macrophages to biomaterials causes activation and induction of this procoagulant activity by inducing the ex-

Tissue Engineering of Prosthetic Vascular Grafts

pression of tissue factor, a transmembrane protein.151 This factor is capable of initiating the extrinsic pathway of coagulation,152 leading to fibrin generation.153 In vitro studies have shown that, in particular, the adherence to Dacron initiates more procoagulant activity than adherence to ePTFE. This seems to relate to the ability of Dacron to adhere more monocytes than ePTFE,154 rather than to its surface characteristics. Hence, the increased deposition of fibrin onto the luminal surface of Dacron grafts over time, as opposed to ePTFE, may relate to the initially higher adherence of monocytes to Dacron. In addition, macrophages are known to secrete both Il-1β and TNF-α following adherence to synthetic materials. These cytokines also enhance tissue factor expression,155 leading to a further induction of PCA and further fibrin deposition. Since Dacron induces higher levels of secretion of these cytokines than ePTFE, this may further facilitate fibrin deposition on the Dacron surface. It would appear, therefore, that differences in the ability to mediate monocyte adhesion and activation and so differentially upregulate procoagulant activity could partially explain the observed differences between the luminal fibrin layer of Dacron and ePTFE. In addition, tissue factor activity has been associated with membrane vesicles that apparently are shed concomitantly with procoagulant activity.156,157 Macrophages in the depth, therefore, could theoretically influence fibrin deposition on the surface by shedding vesicles containing transmembrane tissue factor. The higher number of macrophages present in the depth of Dacron grafts as opposed to ePTFE grafts would, therefore, augment coagulation. In summary, although fibrin forms a pro-angiogenic matrix and seems to be an “ideal” scaffold for wound healing, it may also present its Janus face under different circumstances. It seems likely that the unusual situation of ongoing surface thrombogenicity within the blood stream also produces an unusual fibrin matrix which is rich in fibrinogen and ingrowth inhibitory cytokines.

Macrophages Macrophages dominate the tissue reaction against prosthetic vascular implants more persistently than any other cell type. The deviation from a normal healing process, however, is not their early presence but the tenacity with which they reside in the graft structure even after years of implantation. Their physiological role in early phases of wound healing is opposed by the perseverance of a chronic inflammatory reaction to which they significantly contribute in later stages. Although it is difficult to answer the often recurring question of whether macrophages are friends or foes in prosthetic graft healing, it seems that they are both: friends at the beginning and foes at the end. Their active secretion of cytokines may act as a key chemotactic and mitogenic factor for capillary and connective tissue ingrowth in the early phases of wound healing. Their continual presence, however, may increasingly derail the delicate chronology of cytokines required for the successful accomplishment of healing and also result in material degradation. The difficulty in identifying the key events responsible for this adverse development lies in the multitude of contributing fac-

The Lack of Healing in Conventional Vascular Grafts

tors. On the prosthetic side the different materials, structures and microstructures of the graft are all capable of separately influencing the foreign body reaction against the implant. From the biological standpoint, distinctly different environments characterize the luminal and the outer aspect of grafts. The dynamics of macrophage infiltration can best be studied on ePTFE grafts, with their fairly even wall structure. In the first weeks of implantation one can typically observe a sandwich structure of macrophage populations in the graft wall with a relatively distinct accumulation of Ham 56 and CD68 positive cells at the inside and outside surface, separated by a cell-free fibrin matrix in between.13,18,54,28,33 This triple layered infiltration pattern clearly proves that recruitment occurs from both sides. At the luminal surface the extravasation of blood-borne cells into the porosity of the material occurs simultaneously with insudation of blood proteins immediately following implantation. At the adventitial surface, however, only a few monocytes may extravasate during the initial bleeding phase prior to hemostasis. Subsequently, macrophages have to get there by adhering to capillary endothelial cells and transmigrating through basement membranes via the sequential action of a number of adhesion molecules on both the monocyte (leukocyte specific β2 integrins)158 and the endothelial cell surface (ICAMs).159 It follows, therefore, that macrophages, irrespective of whether they originate from the blood or from transmigration into the adventitial tissue, are faced with a similar task to begin with, namely, to infiltrate a provisional fibrin matrix probably not much different from that found in a normal wound. During infiltration, macrophage migration is facilitated by a low level of basic uPA activity. This should also be sufficient to liberate and activate platelet derived TGF-β from the fibrin matrix160 which was previously incorporated during hemostasis-related fibrinogenesis.139 Since TGF-β upregulates the expression of integrins on the surface of the blood monocytes thereby promoting adhesion and migration,161 it acts as a strong chemotactic agent to recruit macrophages to the inflammatory site.122 Moreover, it also upregulates the uPA activity of macrophages162 to further facilitate migration. Macrophages are also capable of binding and internalizing fibrin through the Mac-1 receptor, which augments fibrinolysis by a plasmin-independent mechanism.163 The degradation products of fibrinolysis, in return, also act as chemoattractants, recruiting still more macrophages into the implant site.164 In addition, trafficking of monocytes to inflammatory sites also involves the superfamily of chemoattractant cytokines (chemokines ) and their receptors.165 Specifically, the C-C chemokine MCP-1 has been demonstrated immunohistochemically to be expressed by macrophages residing at the implant site as early as 48 h postimplantation. 166 This sequence of events illustrates how this initial phase of healing is dominated by self-augmentation of macrophage chemotaxis. Until the macrophage actually contacts the material surface, the recruitment into the wound fibrin of the prosthetic graft does not deviate from that taking place in normal wound healing.

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Following material contact, one of the most important determinants in the chronic inflammatory reaction against synthetic vascular grafts is the biomaterial itself. Adherence, spreading and the secretion of inflammatory mediators like TNF- α, 167-171 IL-1 β 154,163,167,172-177 and IL-6167-169,171 have been extensively investigated on all major biosynthetics. Dacron, for instance, was shown to lead to a higher density of cells with a morphology indicative of activation than ePTFE,154 as well as to higher TNF-α,178 IL-6168 and Il-1β secretion.168,179 Not only does Dacron induce a greater response, but it seems to induce an earlier response as well, with the IL-1 peak occurring on day 4 on Dacron but only on day 7 on all other biomaterials.177 Furthermore, material differences influence not only the ability of macrophages to secrete cytokines but also the expression of other markers of activation, for example, MHC-II,180 Cd11b181 or TGF-β.29 Expanded PTFE and Dacron also differentially induce alteration of expression of subunits of the β2 integrins,182 which mediate both cell-substratum and cellcell interactions. However, there is mounting evidence that these effects are indirect, rather than directly material-mediated. The different abilities of the various materials to adsorb plasma proteins, in particular, appear to be a main determinant behind the distinct macrophage reaction against certain biomaterials.183-185 On materials such as polyethylene terephthalate (PET), polytetrafluoroethylene (PTFE) and polyurethanes, the major blood proteins albumin, fibrinogen and IgG are the predominant proteins adsorbed onto the material surface, with other blood proteins adsorbed to a lesser extent.183-185 Protein adsorption is not dependent, however, on relative protein concentrations in the blood,184 but is determined primarily by the nature of hydrophobic interactions between the material surface and hydrophobic domains of the protein. Since most proteins have a net negative charge, secondary electrostatic interactions are also important in protein-material interactions and are governed by the chemical nature of the material. The overall combination of hydrophobic and electrostatic interactions will have several effects on the process of protein adsorption. Firstly, different biomaterials adsorb different types and relative concentrations of proteins, with more hydrophobic materials adsorbing more protein. It has also been shown that the nature of the adsorbed proteins changes over time with varying rates of adsorption versus desorption on different materials.183 Finally, material-dependent conformational changes as a result of these interactions can be detected by differences in elutability of proteins on different chemical structures.186,187 These changes in conformation differentially expose hydrophilic domains to the aqueous environment, which may or may not contain sites capable of mediating cell adhesion. When comparing two hydrophobic surfaces of different chemical structure, therefore, although a nonpolar hydrophobic material may adsorb more protein, the influence that interactions with polar groups on a chemically different surface will have on the conformation of these adsorbed proteins will ultimately determine the differences in cellular interaction. This may explain why PTFE which, with a surface tension of 18.5 dynes/cm2,188 is certainly more

32

hydrophobic than Dacron (43 dynes/cm2)188 and should thus adsorb more protein,189,190 mediates less cell adhesion. One can imagine that the increased polarity of the carboxyl groups of PET compared to the apolar fluorine groups of PTFE affects the conformation of the adsorbed proteins to the extent that Dacron binds more macrophages than PTFE.154 Of the predominant proteins absorbed, fibrinogen has been shown to be a primary mediator of this adhesion as conformational changes on adsorption expose the P1 epitope which is recognized by the Mac-1 receptor (Cd11b/cd18, 191-194 αmβ2) of the macrophage. This receptor also binds to adsorbed fragments of the complement component C3 (C3bi),195 but studies in complement—depleted mice indicate that this is not a major mechanism of monocyte adherence.191-193 As we have previously noted, the luminal surface of a vascular prosthesis is exposed to a practically unlimited pool of fibrinogen derived from the circulating blood and released by platelets on activation. The insudation of this fibrinogen into the depth of the graft structure facilitates macrophage adhesion to the material. Thus, the superior healing pattern of identical implants in the subcutaneous position, with limited exposure to circulating or platelet fibrinogen, may well be partly explained by this phenomenon. Apart from the material effects on adhesion, the structure of grafts certainly plays another important role with regard to the macrophage response. Increasing diameters of synthetic fibers, for instance, result in increasing numbers of adherent, nonspreading macrophages—a morphology indicative of nonactivated cells.196 This effect may be due to the associated changes in curvature. In addition to microstructure, the overall porosity of the material will also have important implications. Ultimately this will be constrained by two factors: high enough porosity to allow tissue ingrowth, yet low enough to limit continuous insudation of fibrinogen. It seems reasonable to expect a diminished chronic macrophage response in grafts where a temporarily sealing ingrowth matrix limits the ongoing insudation of plasma and hence limits monocyte adhesion to the material through fibrinogen adsorption. As the deposition of fibrin onto the graft surface continues, the inner macrophage zone increasingly becomes wedged between two almost acellular fibrin matrices, one in the graft center and one on the blood surface.2-4,7,60,66,72,82 As discussed under “Fibrin”, the most likely explanation for this mitigated macrophage recruitment from the blood seems to be the changing physical nature of the fibrin matrix of the inner capsule due to different fiber structure and concentrations of fibrinogen. In the outer wall of the graft, however, a macrophage-dominated cell infiltrate continues to populate the outer one third to one half of the prosthetic structure within 1-3 months.27,28 As the macrophage population migrates into the depth of the graft structure, it must eventually overstep the maximum distance of a cell from a capillary and thus get into a mild hypoxic state. Macrophages exposed to hypoxia, however, are synthesizing and releasing increased amounts of PDGF as well as aFGF and bFGF.197 Medium conditioned by hypoxic macrophages in culture is

Tissue Engineering of Prosthetic Vascular Grafts

mitogenic for hypoxic endothelial cells through the release of the FGFs. This suggests a paracrine mechanism in which hypoxia stimulates macrophages to release mitogenic factors for hypoxically preconditioned endothelial cells, thereby promoting angiogenesis. Additionally, TNF-α, which is also upregulated in macrophages by hypoxia,198 is known to play a role in macrophage mediated angiogenesis, presumably through the activation and recruitment of further monocytes.199 The role of TNF-α in angiogenesis is, however, controversial and seems to depend on the biological context. 200,201 These effects are further potentiated by upregulation of MIP-1, a potent macrophage chemoattractant, during oxygen deprivation.198 Therefore, macrophages residing in the outer half of synthetic grafts together with VEGF-releasing infiltrating neutrophils136 are likely to continue to contribute to a pro-angiogenic milieu at this stage. The low concentration of TGF-β released with migration-associated degradation of the provisional wound fibrin should not be inhibitory for endothelial cell proliferation at this stage.137 As the organization of the macrophages within the graft wall changes and transmural endothelial cell ingrowth begins, the situation where the outer wall of the graft is dominated by single macrophages also slowly begins to change by the increasing appearance of foreign body giant cells.13,19,52,55,64 These may persist for the lifetime of the implant. As macrophage adhesion is determined by material characteristics, so is macrophage fusion to form giant cells dependent on material surface as well as on structure. It has been shown that relatively hydrophilic material surfaces induce more fusion than relatively hydrophobic surfaces.202-206 Theoretical analyses show that the time-dependent foreign body giant cell formation is also influenced by the initial density of adherent macrophages.203,204 Therefore, any surface characteristic which influences the number of adherent cells will also influence the extent of giant cell formation. Most conspicuous is the absence of foreign body giant cells (FBGCs) inside the wall structure of ePTFE grafts, compared to Dacron and polyurethane. Although a rim of giant cells often demarcates these prostheses against the surrounding tissue19 and the interstices are sometimes packed with single macrophages, no fusion occurs. We have regularly observed this phenomenon in both high porosity and low porosity ePTFE in the primate. Although the internodal distances of a 60 µm ePTFE graft provide both wide localized ingrowth spaces of approximately 47.94 ± 9.90 µm and a significant adhesion area on the stretched fibrils, macrophages prefer a solitary existence under these circumstances. Similarly, the PTFE nodes—which are the adhesion site for foreign body giant cells on their smooth end on the exterior of the grafts— remain free of giant cells inside the graft wall where numerous fibrils subdivide their surface. The absence of giant cells on the nodal surface subdivided by numerous fibrils indicates that giant cell formation may be suppressed by geometrical space limitations. However, given the in vitro evidence that fewer monocytes adhere to ePTFE and it is less inflammatory than Dacron,168,178,179,182 and that a more hydrophobic surface decreases the propensity of single macrophages to fuse,202-204,206 the inflammatory potential of the

The Lack of Healing in Conventional Vascular Grafts

surface may also play a role in giant cell formation where a more chemically inert, more hydrophobic ePTFE surface reduces fusion. Not only do mononuclear cells gradually fuse to form more giant cells in a material-dependent manner, but further complexity is added by the changing nature of the mononuclear macrophage population over time. Macrophages found in intimate contact with synthetic materials in situ were ED1 and MHCII positive, whereas ED2 positive macrophages were observed lying behind the ED1 positive cells.169 This indicates a pattern of distribution of not only mono—versus multinuclear macrophages, but also immature (ED1 positive) versus mature (ED2 positive), and activated (MHCII positive) versus nonactivated macrophages within the implant site. Theoretical analysis shows that up to 5 weeks of implantation, FBGCs are in fact formed from the fusion of a small percentage of the adherent macrophages which are already present on the third day of implantation.203,204 This raises the issue of fusogenic potential of different subsets of cells. Expression of cell surface receptors may be a prerequisite. For example, inhibitors of mannose receptor activity have been shown to inhibit IL-4 induced fusion in vitro, suggesting that this receptor may play a role in giant cell formation.207 Elaboration of certain cell surface markers in a material or environment dependent manner may be required for a higher propensity for giant cell formation. If it is difficult to decide whether macrophages are beneficial for graft healing or not, it is even more difficult to decide whether the presence of foreign body giant cells represents a state of pacification or a particularly detrimental variant of chronic inflammation. One indication for the FBGCs role as a mediator of pacification of inflammation arises from the fact that their fusion can be induced by IL-4.99,208 IL-4 is secreted by Th2 lymphocytes and is known to downregulate several macrophage inflammatory functions, for example, the release of certain inflammatory mediators,209 the secretion of reactive oxygen intermediates210 and collagenase production,211 to mention a few. In this context it has been proposed that giant cells may represent a attempt to “mop-up” the deleterious effect of persistent chronic foreign body reaction. In contrast, a detrimental role for the giant cell is suggested by the fact that material surface cracking occurs directly beneath adherent foreign body giant cells, implicating the secretory products of these cells in the degradation of biosynthetics.212 The presence of actin-containing adhesive structures202 indicates that the adhesion of the giant cell to the biomaterial forms a sealed compartment with a high local concentration of degradative enzymes. It is also peculiar that the accumulation of giant cells in Dacron regularly coincides with the development of huge convolutes of capillary sinuses in the vicinity. These widely dilated and irregularly shaped endothelial sacs are principally devoid of any second cell layer (Figs. 1.11, 1.18). In contrast, 60 µm ePTFE, with its scanty presence of foreign body giant cells, shows arterioles in its outer interstitial spaces with one or more layers of smooth muscle cells. This histological observation, where the lack of giant cells correlates with the absence of malformed vessels in the form of

33

capillary sinuses and the presence of smooth muscle cells, seems to indicate that the giant cell contributes to the misdirecting of the process of transmural capillary ingrowth. In situ hybridization of human explants has demonstrated that interstitial macrophages and foreign body giant cells are still secreting Il-1β even after numbers of years.213 Although Il-1β is known to stimulate both smooth muscle cell proliferation and fibronectin deposition in atherosclerotic plaque formation, 214 it has been suggested that this interleukin can also inhibit vascular smooth muscle proliferation in an autocrine fashion through nitric oxide upregulation.215 The giant cells, therefore, may be responsible for not only affecting capillarization but also inhibiting smooth muscle cell ingrowth through persistent Il-1β secretion. In summary, the true character of the FBGC remains elusive. On the one hand, their persistence at the implant site and their destructive role in biodegradation, as well as their possible role in the mitigation of vessel formation, seem to label them as villains in incomplete graft healing. Yet, the evidence supporting a role for a typically antiinflammatory cytokine—IL-4—in stimulating giant cell formation is fairly convincing both in vitro and in vivo.99,208 Given that inflammation associated with normal wound healing is resolved because of the transient nature of the macrophage response, it would appear that the inability to remove the inflammatory stimulus (a nondegradable biomaterial) confounds even the best intentions of these cells to mediate healing. The consequence is a chronic ongoing inflammation and the derailing of the normal sequence of healing events. In summary, two distinct zones develop in prosthetic vascular grafts over time: one close to the blood surface which becomes increasingly uninhabitable for macrophages, and one in the outer portion of the graft which is densely packed with macrophages and later on with foreign body giant cells. It is well documented that macrophages infiltrating a wound environment stimulate the proliferation of endothelial cells and, hence, angiogenesis. However, the apparent lack of correlation between the number of resident macrophages in the interstices of grafts and the extent of transmurally ingrowing vessels indicates that macrophages are not alone in affecting endothelial ingrowth, but that other factors, for example, the nature of the ingrowth matrix, may affect the overall extent of capillarization. Aside from an often impenetrable fibrin matrix, the protracted ingrowth of connective tissue, specifically smooth muscle cells from outside, may well have to do with ongoing macrophage activation in the presence of persistent material stimulus. The secretion of inhibitory cytokines like Il-1β by both macrophages and foreign body giant cells may be partly responsible. What is becoming increasingly obvious is that a macrophage is not always a macrophage and that the contribution of each of these subsets of macrophage populations to the lack of healing of the vascular graft remains to be investigated. Overall, the macrophage remains the Jekyll and Hyde of graft healing, but as our insight into the mechanisms of healing grows so does our understanding of which critical questions remain to be answered in the elucidation of the role of this enigmatous cell, namely:

34

Tissue Engineering of Prosthetic Vascular Grafts

1. If monocyte adhesion to biomaterials is prevented completely, do we compromise healing in any way; 2. To what extent does engineering tissue ingrowth modulate the deleterious effect of the inflammatory cell by promoting selective cellular infiltration before the development of chronic inflammation; and 3. Can we identify secretory products of macrophages or giant cells which specifically mediate the mitigation of tissue ingrowth, either endothelial sprouts or smooth muscle cells?

Prosthetic Wall Vascularization The “restitutio ad integrum” healing of a vascular prosthesis—the formation of a neo-artery in place of a removed segment—would necessitate the formation of vasa vasorum in the graft wall, accompanied by the growth of a fully functional medium in an inflammation-free environment. The second best solution—the formation of granulation tissue with the resulting conclusion of the healing process through scar tissue—would at one stage also require extensive capillarization. Even if conventional synthetic grafts are lacking the sophisticated biological information necessary for “ad integrum” healing, their porosity would in many cases at least allow the ingrowth of granulation tissue. From our daily surgical practice we are all familiar with the alacrity with which the formation of such granulation tissue and the subsequent stabilization through scar tissue is usually accomplished. The tissue gap in technically the best of all surgical wound closures is often more than one millimeter wide, but capillarization is not only complete, but scar tissue formation is already beginning, as early as after seven days. It is therefore surprising that an equally short distance in the case of the wall of prosthetic vascular grafts does not accomplish the same degree of full thickness capillarization even after years of implantation. In order to further elucidate this major shortcoming in graft healing. one needs to ask two key questions: 1. Why does adventitial capillary sprouting into the outer fibrinous capsule and later on into the interstices of the prosthetic graft occur at such a slow pace and eventually come to a halt; and 2. Why is anastomotic endothelial cell outgrowth so futile concerning both surface endothelialization and capillary sprouting into the depth? The initial angiogenic response of the surrounding tissue to a freshly implanted vascular prosthesis does not seem to differ from any other wound repair. This early phase is typically characterized by infiltration into the wound fibrin of a relatively high proportion of polymorphonuclear granulocytes, but also lymphocytes and a few macrophages. Capillaries within the adventitia sprout into the freshly formed fibrin matrix within days of injury.2,13,48,62,63,76,94 During this formation of early granulation tissue, the matrix is probably inconspicuous with regard to fibrinogen concentration, fiber constitution and cytokine content and, as such, is naturally pro-angiogenic. The sprouting endothelial cells bind via αvβ3 to the RGD site of the fibrin α-chain and migrate into the pores of the graft. This also promotes survival of the angiogenic sprout by sustaining an antiapoptotic response through inhibition of p53 activation.216

Apart from the favorable matrix environment, white blood cells provide various pro-angiogenic stimuli at this stage. Infiltrating neutrophils are a source of VEGF,136 and macrophages migrating into a presumably mildly hypoxic environment secrete acidic and basic FGF as well as PDGF.197 These cytokines are both chemotactic and mitogenic for endothelial cells. Once again, integrin mediated binding to the wound matrix is central to these responses, with both αvβ3 and αvβ5 involved in growth factor-stimulated angiogenesis, although via different mechanisms:59 bFGF-induced angiogenesis is mediated by the αvβ3 integrin, whereas αvβ5 integrin is crucial in VEGF-stimulated angiogenesis. Although the initial phase of vascularization resembles normal wound healing, the subsequent phase of centripetal vessel ingrowth through the interstices of the graft is already substantially slowed and often only reaches the outer half of the graft wall.60,78 Considering the fact that capillarization of several millimeters of wound fibrin is normally accomplished in humans within days, the time period of several months required for achieving this ingrowth can only be explained by a significant deviation from normal wound healing. One obvious difference between normal and graft healing is the nature of the chronic inflammatory response. Specifically, the macrophage-derived giant cells which gradually build up in the outer parts of the graft27,28,33,60,62 are only formed in response to a foreign body. In high porosity Dacron grafts, this foreign body reaction is not only associated with mitigated ingrowth but also the formation of hugely dilated capillary sacs as opposed to regular vessels (Figs. 1.11, 1.18). Moreover, in some models, the conspicuous lack of smooth muscle cells, and therefore of vessels other than capillaries, indicates that not only angiogenesis but also musculogenesis is inhibited. Given these unique circumstances, it should be possible to describe biological interactions, which at this stage of graft healing, may be responsible for the gradual mitigation of transmural ingrowth. Histological analysis suggests that probably both matrix- and cell-mediated phenomena are responsible for the inhibition of angio- and musculogenesis. As far as the matrix is concerned, we have already discussed the possibility of a gradual transition in the nature of the fibrin from the outer wound to a fibrinogen-rich blood surface which may gradually slow down the capillarization from outside to inside. In this regard, it is interesting that studies on matrix control of angiogenesis217,218 have observed that it is specifically capillary morphogenesis that depends on extracellular matrix (ECM).219 Especially, the ability of endothelial cells to exert mechanical forces on the surrounding matrix,217 and the subsequent development of matrix networks218 to guide capillary sprout formation, are matrix determined. Physical and/or biochemical aspects of fibrin formation will, therefore, affect the malleability of the matrix and hence the susceptibility to cell mediated traction and the realignment of matrix fibers. Hence, the significance of the qualitative change in the matrix towards the lumen may be inhibition of capillary ingrowth. In addition to matrix mediated effects, two histological observations hint at cellular events playing a role in the mitigation of vascularization. Firstly, there is the fact that capillary sprouting increasingly comes to a halt towards the

The Lack of Healing in Conventional Vascular Grafts

acellular inner fibrinous capsule, while the macrophagedominated chronic inflammatory tissue, which is limited to the outer half of the graft, reaches a state of sufficient vascularization.2,11,27-29,60 It would appear that, although hypoxic macrophages may initially be responsible for a pro-angiogenic stimulus, the vascularization which has taken place in the meantime in the outer half of the graft re-establishes a normoxic environment. With an increasing oxygen tension in the tissue, one would expect a downregulation of the secretion of certain endothelial cell mitogens220,221 and hence a downregulation of the angiogenesis. Moreover, the shortlived neutrophil response may further deprive angiogenesis of a potent stimulator—VEGF.56,136,222 The second and perhaps most unique observation is not that angiogenesis gradually slows down, but the fact that vessel formation is so derailed that the endothelial cells form sinuses as opposed to tubes. It seems likely that this sinus formation may be due to the lack of appropriate intercellular alignment and adhesion events, which are as critical to the process of angiogenesis as the migration and proliferation of endothelial cells.218 One hint of what the mechanism may be comes from an in vitro model of bFGF-induced angiogenesis. It was observed that, in fibrin gels, stimulation of endothelial cells with bFGF alone resulted in large lumen sinuses similar to those seen in grafts, whereas the addition of TGF-β at picogram concentrations reduced the lumen dimensions.223 Addition of nanograms of TGF-β mitigated lumen formation completely. Given these in vitro observations, it would appear that ultimately the 3-dimensional organization of capillary sprouts depends on a balance of proliferating versus differentiating stimuli, where an overriding mitogenic stimulus (in this model, bFGF) results in large sinuses but increasing concentrations of a differentiating stimulus (TGF-β) arrests proliferation and so facilitates tube formation. Driven to the other extreme, an overriding differentiating stimulus arrests both proliferation and tube formation. If one concludes that the sinuses are the result of overriding mitogenic signals in the outer zone of the graft wall, where could these signals originate? As previously mentioned, histologically we observe that sinus formation is associated with the gradual buildup of giant cells. The obvious association of these cells with the capillary sinuses indicates that they may be secreting an abundance of endothelial cell mitogens. Although macrophages themselves are known to stimulate endothelial proliferation,197,220 the limited information available on the nature of secretory products of the foreign body giant cell makes it difficult to define exactly what paracrine mechanism may be involved. To further complicate the situation, the success of graft wall vascularization appears to depend not only on the interaction of these pro- and anti-angiogenic influences of inflammatory cells and matrix, but also on the interplay of these two factors with the angiogenic vigor of host endothelial cells. The example of 60 µm ePTFE grafts in different animal models illustrates best that there seems to be a net effect between the endothelial cell’s inherent ability to rapidly proliferate and migrate on the one hand and the gradual build up of an environment which inhibits angiogenesis on the other. In the juvenile animal model of the yellow ba-

35

boon, for instance, capillary outgrowth is rapid and thus completed before a changing environment can inhibit it.30,32,34 Under senescent circumstances as in humans78 and the chacma baboon, however, the protracted angiogenic potential of endothelial cells does not allow completion of capillary ingrowth prior to the change in the biological environment. This seems to explain why the same graft leads to rapid spontaneous endothelialization in the one model but disproportionately fails to achieve this goal in another. While unfavorable matrix characteristics, continual macrophage presence and host specific endothelial cell characteristics offer a possible explanation for the slowing down of capillary ingrowth, the lack of musculogenesis within the graft wall is more challenging to explain. With potent tools of tissue engineering available, one needs to overcome the deep seated suspicion of the presence of smooth muscle cells based on their role in intimal hyperplasia. This suspicion prevails despite reports that smooth muscle cell proliferation reaches an equilibrium on prosthetic vascular grafts,61,93 depending partly on functional interactions with physiological regulators like blood flow.54 However, the goal of eventually emulating functional arteries through “smart” engineered scaffolds will necessarily need to assign the significant role of creating a “healthy” neo-media to smooth muscle cells. It is therefore paramount to try to understand mechanisms of ingrowth inhibitions for smooth muscle cells also and not only for endothelial cells. Therefore, one must see two peculiar observations in prosthetic midgraft healing in this context: the absence of smooth muscle cells in most parts of the graft wall2,16,26,27,94,224 and at the same time the presence of smooth muscle cells underneath luminal endothelial cells,2,14,30-33,35,59,67,77,81 sometimes separated by acellular fibrin59 from transmurally ingrowing connective tissue. With transdifferentiation of fibroblasts being only a remote possibility, the presence of smooth muscle cells seems to be dependent on the differentiation of mural/mesenchymal cells in the wake of angiogenesis.106,107 Embryonic data suggests that endothelial tube structures form first and direct the subsequent recruitment of the surrounding cells— pericytes in small vessels and smooth muscle cells (SMC) in larger vessels. This promotion of mural cell proliferation and migration is primarily achieved through the synthesis and secretion of platelet derived growth factor-BB (PDGF-BB) by endothelial cells, which upregulates smooth muscle cell expression of collagenase, αvβ3 and PDGF-receptors. The upregulation of these factors allows the mural cells to move toward the endothelial cell, eventually establishing intercellular contact through adherens junctions. This intimate contact between the two cell types drives the expression and activation of TGF-β, resulting in a local high concentration sufficient to inhibit endothelial cell migration and proliferation225 as well as to promote differentiation of the pericytes/smooth muscle cells. This delicate cytokine balance between endothelial cells and smooth muscle cells reflects the interdependence of these two cell types, further emphasizing that smooth muscle cells require the presence of endothelial cells for migrating into tissue. If, on the one hand, capillaries eventually reach the inner fibrinous capsule of prosthetic grafts and, on the other

36

hand, mesenchymal cell differentiation into smooth muscle cells normally “follows” the outgrowing endothelial sprouts, why are capillaries often found in abundance but without accompanying smooth muscle cells? Similarly to the inhibition of capillarization, both cell-matrix and cell-cell interactions seem to play a role in the inhibition of musculogenesis. As discussed under “Fibrin”, platelet activation at the blood surface may result in relatively high TGF-β concentrations in the nonfavorable fibrin matrix. Should migration-associated degradation liberate this TGF-β, it may indeed prematurely inhibit proliferative responses of smooth muscle cells and switch the phenotype of these cells to a resting one. In this way, the biological signals encoding the maturation and eventually the arrest of a proliferative event may be prematurely transmitted to the smooth muscle cells as a result of the changing fibrin environment over time. However, as tempting as TGF-β may be as an explanation for the mitigation of further capillary ingrowth into the inner fibrinous capsule, it does not explain the absence of smooth muscle cells in areas of dilated endothelial sinuses towards the adventitial surface where the fibrin is assumed to be closer to normal wound fibrin both in structure and cytokine content. Because these sinuses are associated with macrophages and giant cell formation, one explanation may be the immunohistochemical observation that macrophages and giant cells are secreting Il-1β even after years of implantation.213 This cytokine has been show to upregulate smooth muscle cell nitric oxide production215 and could, as such, negatively regulate smooth muscle cell mitogenesis in areas of macrophage infiltration, in contrast to its role in the pathogenesis of atherosclerosis.214 More likely, however, is an absence of PDGF-BB secretion by the endothelial cells, because they never reach maturation associated with vessel formation. In addition, we have already discussed the role of a slowly changing matrix on the mitigation of full thickness wall capillarization. Such arrested endothelial cells may cease to send out chemotactic signals to smooth muscle cells in form of PDGF. Interestingly, although ingrowing capillaries in high porosity Dacron grafts were completely devoid of smooth muscle cells in the chacma baboon model, 60 µm ePTFE grafts in chacma and yellow baboons showed arterioles with layers of actin positive smooth muscle cells (Fig. 1.9). This indicates that in those grafts in which capillarization was able to overcome ingrowth inhibition at an early stage and proliferation of the endothelial cells was not arrested, the smooth muscle cell component is normally developed.2,30-33,35,46,77,81 This supports the hypothesis that mitigation of capillarization influences smooth muscle cell migration. Moreover, it is noticeable that the combination of high porosity ePTFE with a juvenile and vigorously endothelializing animal model achieves the most differentiated layers of smooth muscle cells.30-33,35 These smooth muscle cells are observed in immediate proximity to endothelial cells within the initial 2-3 weeks of implantation when biological circumstances seem to be more in favor of angiogenesis, and hence smooth muscle cell recruitment, than at a later stage. Therefore, if the fibrinogenic potential of a graft is relatively low, the angiogenic vigor of endothelial cells of

Tissue Engineering of Prosthetic Vascular Grafts

a species high and the porosity favorable to allow ingrowth together with early sealing of the interstices against continual fibrin/fibrinogen insudation, capillaries may reach across the entire distance before the biological environment changes. As a consequence, no premature maturation event would prevent the endothelial cells from chemotactically attracting mural cells to trail them. This scenario seems to explain best why 60 µm ePTFE grafts in the juvenile yellow baboon show such a distinctly different healing pattern with regard to smooth muscle cell ingrowth compared to other grafts and species. However, if the fibrinogenic potential of a graft is relatively high and the angiogenic potential of the host endothelial cell low, it is reasonable to imagine that the slowly changing biological environment allows the first part of a chronological sequence but prevents the second one. As the circumstances deteriorate to a point where the entire process begins to halt, the first step—in this case, capillary sprouting—has already materialized while the subsequent step—the site-directed migration of mural cells—has not even commenced yet. When both endothelial cells and smooth muscle cells do reach the blood surface, additional large vessel signals appear to play a role in their relationship. Flow conditions and compliance mismatch may result in an overriding stimulus for smooth muscle cell proliferation229 despite their direct contact with endothelial cells, which would normally terminate smooth muscle cell mitogenesis through the effect of TGF-β. It is well established that the endothelium transduces flow-related signals and directs smooth muscle cell proliferation when required. It is not surprising, therefore, that expression of the smooth muscle cell mitogen PDGF-AA has been detected in the endothelial cells of the well developed intima of grafts implanted in the yellow baboon,52,226 while very little PDGF-B chain is detected. Considering the complexity of, for example, PDGF signaling where PDGF is a homo- or heterodimer of two different chains with three resulting isoforms, not to mention two different receptors selectively binding all or only one of these isoforms, it is possible that these factors may combine to generate a wide variety of results and afford the cell a relatively tight control of the outcome.227,229 This would explain why PDGF-AA is present during intimal proliferation although PDGF-BB normally directs smooth muscle cell migration and proliferation during angiogenesis. If premature maturation signals are the reason for the stoppage of transmural ingrowth of capillaries at the inside of the inner fibrinous capsule, is it possible to apply the same interpretation to the ingrowth stoppage of anastomotic surface endothelium? At first thought, events only partially resemble those occurring in the graft wall. Since angiogenesis is a complex process including matrix degradation, migration and proliferation, the earlier termination of endothelial outgrowth in the wall matrix may be due to the inhibition of either of these three. Since surface migration does not require matrix degradation to the same extent, it must be either proliferation or migration which is primarily affected under these circumstances. Extrapolation of the number of population doublings230 required to cover the transanastomotic ingrowth area clarifies that it is not the divid-

The Lack of Healing in Conventional Vascular Grafts

ing potential of the endothelial cell which is the limiting factor of endothelial outgrowth. Therefore, inhibitory factors other than an exhausted mitotic capacity must play a role. One explanation could be activated platelets that accumulate in and on the pseudointimal layer throughout the length of the graft. The complex signaling arising from platelet degranulation and the high local TGF-β concentration which results could affect endothelial cell proliferation at the surface. However, there must be an additional explanation for transanastomotic ingrowth stoppage, because it is also seen in grafts with minimal surface fibrin or platelet coverage. The possible involvement of an additional factor becomes apparent in Dacron grafts, where the outgrowth of endothelium is more pronounced in a high porosity (knitted) Dacron40,46,47,60,68,81,107 than in a low porosity (woven) one.40,43 The difference between the two types of Dacron grafts obviously lies not in the composition of the fibrinous pseudointima but in the ingrown tissue underlying this fibrin matrix. Could this ingrowing tissue in some way be facilitating endothelial ingrowth from the anastomosis so that the gradual cessation of transmural ingrowth becomes reflected in a similar mitigation of outgrowth from the anastomoses? Paracrine signaling over such distances is known to occur during embryogenesis, so it is not beyond probability that ingrowing tissue directs transanastomotic outgrowth but the underlying mechanism has yet to be determined. In summary, the most striking shortcomings of prosthetic wall vascularization, namely, the failure of the capillaries to reach the luminal surface as well as the absence of mature smooth muscle cells and hence arterioles, could be partly attributed to both the surrounding matrix and the surrounding inflammatory cells. Graft healing therefore seems to be a race between the endothelial cell’s angiogenic potential and the buildup of inhibitory influences. Since the endothelial cell remains a relative constant in human implants it makes sense that graft designs should aim at, firstly, preventing the deposition of unfavorable, compacted fibrin and, secondly, attenuating the chronic inflammatory response. This could well be achieved by incorporating an ingrowth matrix which not only temporarily seals the porosity of the graft wall to protein insudation but at the same time selectively promotes vascular cell infiltration while inhibiting the buildup of a prolonged macrophage response.

Conclusion Healing of prosthetic vascular grafts implies transmural tissue ingrowth. During the past three decades, however, the focus of research was almost entirely on transanastomotic rather than transmural ingrowth. In spite of this one-sided approach, an analytic effort combining historical data and personal experience with biological principles helped to clarify some main issues involved in the mitigation of the entire healing process of contemporary arterial prostheses. The complete surface endothelialization through transanastomotic ingrowth in animal models under shorter term experimental conditions often suggested that these models differ in principle from the human situation with its stoppage of endothelial ingrowth. In contrast to this

37

widely held opinion, it would appear that in fact transanastomotic endothelialization progressively slows down in all models, although at different rates. One can therefore not extrapolate short term observations to a long term situation. Since the cessation of tissue outgrowth cannot be due to an exhausted mitotic potential of endothelial cells, migration and proliferation appear to get actively arrested by changing biological circumstances. The fibrinogen and platelet-rich environment at the blood surface makes the buildup of a dense and fine fibrillar fibrin capsule likely, containing unusually high concentrations of TGF-β. Such a matrix seems theoretically capable of prematurely turning the outgrowing endothelium into a noncycling, resting intimal tissue. Porosity has always been regarded as a crucial determinant for tissue ingrowth. However, porosity alone is not sufficient to quantify the available space for ingrowth. Even in so-called high porosity ePTFE grafts for instance, fullthickness transmural ingrowth of microarterioles is only possible in a marginal percentage of interstitial spaces. The graft wall structure may provide the simplest level of inhibition, and therefore this factor needs to be considered. The dimensions of the available ingrowth space in ePTFE grafts have not previously been assessed in order to determine the dimensions of a transmural channel. The mathematical calculations presented for these grafts in this chapter reveal that porosity was previously overestimated as a sole determinant for tissue ingrowth. Similarly, the measurements of interfiber spaces in Dacron grafts revealed a maximum space sufficient for capillaries, and limited access to arteriolar ingrowth. On the biological level, a continual buildup of seemingly adverse conditions eventually prevents the completion of transmural tissue ingrowth in most of the cases. In those instances in which microvessels successfully reach the blood surface, it appears that tissue ingrowth is completed more rapidly and hence before the environment turns inhibitory. When identifying inhibitory structures, the compact inner fibrin capsule and the outer zone, with massive accumulation of foreign body giant cells, appear to be the primary culprits. The special composition of the internal fibrinous capsule provides a reasonable explanation for the fact that it is almost impenetrable for capillaries. In contrast, the outer graft zone which is packed with foreign body giant cells seems rather to affect the quality of angiogenesis by derailing the delicate process of endothelial—and smooth muscle cell— sprouting. While unrivaled endothelial proliferation leads to the formation of hugely dilated capillary sacs as opposed to vessels, no smooth muscle cells follow the angiogenic efforts of the endothelial cells. This lack of musculogenesis occurring in the graft wall corresponds conspicuously with the presence of foreign body giant cells. The presence of giant cells is in turn a consequence of a pronounced macrophage attraction by the synthetic materials. This macrophage infiltration is augmented by fibrinogen insudation from the blood stream. If one tries to draw any conclusions from these observations for the tissue engineering of future grafts the following main guidelines appear to be crucial:

38

• In order to design future experiments with a view towards transmural rather than transanastomotic ingrowth, minimal graft lengths of 15 cm should be observed; • Any structural design of a graft scaffold should make provision for continual transmural ingrowth spaces capable of accommodating at least arterioles in order to theoretically allow the ingrowth of both endothelial cells and smooth muscle cells; • Considering that tissue ingrowth faces an increasingly adverse microenvironment, a key issue to successful healing will be to facilitate rapid completion of vascularization; • Since fibrinogen immobilization to the biomaterial is a main mediator of macrophage recruitment, a temporarily sealing ingrowth matrix may prevent this event and thus reduce the macrophage presence. In summary, we eventually begin to understand key aspects of the aborted healing events in yesterday’s prostheses. This does not only prepare us better for the de novo design of tomorrow’s tissue engineered grafts, but also helps us to change the angle from which we approach graft design: The main emphasis will need to be on the delicate manipulation of biological chronologies in the context of an accurately modeled synthetic scaffold designed to limit inflammatory responses to material contact. References 1. Graham LM, Burkel W E, Ford JW, Vinter DW, Kahn RH, Stanley JC. Expanded polytetrafluoroethylene vascular prostheses seeded with enzymatically derived and cultured canine endothelial cells. Surgery 1982; 91:550-559. 2. Herring MB, Baughman S, Glover J, Kesler K, Jesseph J, Campbell J, Dilley R, Evan A, Gardner A. Endothelial seeding of dacron and polytetrafluoroethylene grafts: The cellular events of healing. Surgery 1984; 96:745-754. 3. Boyd KL, Schmidt SP, Pippert TR, Sharp WV. Endothelial cell seeding of ULTI carbon-coated small diameter PTFE vascular grafts. ASAIO Transactions 1987; 33:631-635. 4. Kao WJ, McNally AK, Hiltner A, Anderson JM. Role for IL-4 in FBGC formation on a poly(etherurethane urea) in vivo. J Biomat Res 1995; 29:1267-1275. 5. Zilla P, Preiss P, Groscurth P, Rösemeier F, Deutsch M, Odell J, Heidinger C, Fasol R, von Oppell U. In vitrolined endothelium: Initial integrity and ultrastructural events. Surgery 1994; 116:524-534. 6. Friedman EW, Hamilton AJ. Polytetrafluorethylene grafts in the peripheral venous circulation of rabbits. The American J Surg 1983; 146:355-359. 7. Zamora JL, Navarro LT, Ives CL, Weilbaecher DG, Gao ZR, Noon GP. Seeding of arteriovenous prostheses with homologous endothelium. J Vasc Surg 1986; 3:860-866. 8. Binns RL, Ku DN, Stewart MT, Ansley JP, Coyle KA. Optimal graft diameter: Effect of wall shear stress on vascular healing. J Vasc Surg 1989;10:326-337. 9. Golden MA, Hanson SR, Kirkman TR, Schneider PA, Clowes AW. Healing of polytetrafluoroethylene arterial grafts is influenced by graft porosity. Journal of Vascular Surgery 1990;11:838-845. 10. Lewis DA, Lowell RC, Cambria RA, Roche PC, Gloviczki P, Miller VM. Production of endothelium-derived factors

Tissue Engineering of Prosthetic Vascular Grafts from sodden expanded polytetrafluorethylene grafts. J Vasc Surg 1997; 25:187-197. 11. Sottiurai VS, Yao JST, Flinn WR, Batson RC. Intimal hyperplasia and neointima: An ultrastructural analysis of thrombosed grafts in humans. Surgery 1983; 93: 809-817. 12. Stumb MM, Jordan GL, DeBakey ME. Endothelium growth from circulating blood on isolated intravascular Dacron hub. Amer J Path 1963; 43:361-368. 13. van der Lei B, Wildevuur RH. Microvascular polytetrafluorethylene prostheses: The cellular events of healing and prostacyclin production. Plastic and Reconstructive Surgery 1988; 81:735-741. 14. Bartels HL, van der Lei B, Robinson PH. Prosthetic microvenous grafting in the rat femoral vein. Laboratory Animals 1993; 27:47-54. 15. Guidoin RG, King M, Marois M, Martin L, Marceau D. New polyester arterial prostheses from Great Britain: An in vitro and in vivo evaluation. Ann Biom Engin 1986; 14:351-367. 16. Stronck JWS, van der Lei B, Wildevuur CRH. Improved healing of small-caliber polytetrafluoroethylene vascular prostheses by increased hydrophilicity and by enlarged fibril length. An experimental study in rats. J Thorac Cardiovasc Surg 1992; 103:146-152. 17. Van Der Lei B, Stronck JW, Wildevuur RH. Enhanced healing of 30 µm Gore-tex PTFE microarterial prostheses by alcohol-pretreatment. British Journal of Plastic Surgery 1991; 44:428-433. 18. Bellón JM, Buján J, Contreras LA, Jurado F. Similarity in behaviour of polytetrafluorethylene (ePTFE) prostheses implanted into different interfaces. J Biomed Mat Res 1996; 31:1-9. 19. Clowes AW, Gown AM, Hanson SR, Reidy MA. Mechanisms of arterial graft failure: I. Role of cellular proliferation in early healing of PTFE prostheses. Amer J of Path 1985; 118:43-54. 20. Florian A, Cohn LH, Dammin GJ, Collins JJ. Small vessel replacement with Gore-tex (expanded polytetrafluoroethylene). Arch Surg 1976; 111:267-270. 21. Hanel KC, McCabe C, Abbott WM, Fallon J, Megerman J. A biomechanical, scanning electron, and light microscopic evaluation Ann Surg 1982; 195:456-463. 22. Mathisen SR, Wu H, Sauvage LR, Usui Y, Walker MW. An experimental study of eight current arterial prostheses. J Vasc Surg 1986; 4:33-41. 23. Sterpetti AV, Lepidi S, Cucina A, Patrizi AL, Palumbo R, Taranta A, Stipa F, Cavallaro A, Santoro-D’Angelo L, Stipa S. Growth factor production after polytetrafluoroethylene and vein arterial grafting: An experimental study. J Vasc Surg 1996; 23:453-460. 24. Campbell CD, Goldfarb D, Roe R. A small arterial substitute: Expanded microporous polytetrafluoroethylene: Patency versus porosity. Ann Surg 1975; 182:138-143. 25. Kenney DA, Tu R, Peterson RC. Evaluation of compliant and noncompliant PTFE vascular prostheses. ASAIO Transactions 1988; 34:661-663. 26. Matsumoto H, Hasegawa T, Fuse MD, Yamamoto M, Saigusa M. A new vascular prosthesis for a small caliber artery. Surgery 1973; 74:519-523. 27. Douville EC, Kempczinski RF, Birinyi LK, Ramalanjaona GR. Impact of endothelial cell seeding on long-term patency and subendothelial proliferation in a small-caliber highly porous polytetrafluoroethylene graft. J Vasc Surg 1987; 5:544-550.

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The Lack of Healing in Conventional Vascular Grafts 170. Hunt JA, Flanagan, BF, McLaughlin PJ, Strickland I, Williams DF. Effect of biomaterial surface charge on the inflammatory response: Evaluation of cellular infiltration and TNF-β production. J Biomat Res 1996; 31:139-144. 171. Yun JK, DeFife, K, Colton E, Stack S, Azeez A, Cahalan L, Verhoeven M, Cahalan P, Anderson JM. Human monocyte/macrophage adhesion and cytokine production on surface modified poly(tetrafluoroethylene/hexafluoropropylene) polymers with and without protein preadsorption. J Biomat Res 1995; 29:257-268. 172. Bonfield TL, Colton E, Anderson JM. Plasma protein adsorbed biomedical polymers: Activation of human monocytes and induction of IL-1. J Biomat Res 1989; 23:535-548. 173. Bonfield TL, Anderson JM. Functional versus quantitative comparison of Il-1β from monocytes/macrophages on biomedical polymers. J Biomat Res 1993; 27:1195-1199. 174. Krause TJ, Robertson FM, Liesch JB, Wasserman AJ, Greco RS. Differential production interleukin 1 on the surface of biomaterials. Arch Surg 1990; 125:1158-1160. 175. Miller KM, Anderson JM. Human monocyte/macrophage activation and IL-1 generation by biomedical polymers. J Biomat Res 1988; 22:713-731. 176. Miller KM, Anderson JM. In vitro stimulation of fibroblast activity by factors generated from human monocytes activated by biomedical polymers. J Biomat Res 1989; 23:911-930. 177. Miller KM, Rose-Caprara V, Anderson JM. Generation of IL-1 like activity in response to biomedical polymer implants: A comparison of in vitro and in vivo models. J Biomat Res 1989; 23:1007-1026. 178. Swartbol P, Truedsson L, Pärsson, Norgren L. Tumor necrosis factor-α and interleukin-6 release from white blood cells induced by different graft maerials in vitro are affected by pentoxifylline and iloprost. J Biomed Mater Res 1997; 36:400-406. 179. Bonfield TL, Colton E, Anderson JM. Protein adsorption of biomedical polymers influences activated monocytes to produce fibroblast stimulating factors. J Biomat Res 1992; 26:457-465. 180. Petillo O, Peluso G, Ambrosio, L, Nicolais L, Kao WJ, Anderson JM. In vivo induction of macrophage Ia antigen (MHC class II) expression by biomedical polymers in the cage implant system. J Biomat Res 1994; 28:635646. 181. Gemmell CH, Black JP, Yeo EL, Sefton MV. Material induced up-regulation of leukocyte Cd11b during whole blood contact: Material differences and a role for complement. J Biomat Res 1996; 32:29-35. 182. Swartbol P, Truedsson L, Pärsson, Norgren L. Surface adhesion molecule expression on human blood cells induced by vascular graft materials in vitro. J Biomat Res 1996; 32:669-676. 183. Anderson JM, Ziats NP, Azeez A, Brunstedt MR, Stack S, Bonfield TL. Protein adsorption and macrophage activation on PDMS and silicone rubber. J of Biomat Sci Polymer Edition 1995; 7:159-169. 184. Anderson JM, Bonfield TL, Ziats NP.Protein adsorption and cellular adhesion and activation on biomedical polymers. Biomaterials 1990; 13:375-382. 185. Ziats NP, Pankowsky DA, Tierney BP, Ratnoff OD, Anderson JM. Absorption of Hageman factor and other human plasma proteins to biomedical polymers. J Lab Clin Med 1990; 116:687-96.

43 186. Rapoza RJ, Horbett TA. Postadsorptive transitions in fibrinogen: Influence of polymer properties. J Biomat Res 1990; 24(10):1263-1287. 187. Slack SM, Horbett TA. Changes in fibrinogen adsorbed to segmented polyurethanes and hydroxyethylmethacrylate-ethylmethacrylate copolymers. J Biomat Res 1992; 26(12):1633-1649. 188. Zisman WA. Ind Eng Chem 1963; 55:18-25. 189. Boffa GA, Lucien N, Faure A, Boffa MC. J Biomat Res 1977; 11:3-17. 190. Noishiki Y. Application of immunoperoxidase method to electron microscopic observation of plasma protein on polymer surface. J Biomed Mater Res 1982; 16:359-67. 191. Tang L, Eaton JW. Fibrin(ogen) mediates acute inflammatory responses to biomaterials. J Exper Med 1993; 178:2147-2156. 192. Tang L, Lucas AH, Eaton JW. Inflammatory responses to implanted polymeric biomaterials: Role of surface-absorbed IgG. J Lab Clin Med 1993; 122:292-300. 193. Tang L, Eaton JW. Inflammatory responses to biomaterials. Amer J Clin Path 1995; 103:466-471. 194. Tang L, Ugarova TP, Plow EF, Eaton JW. Molecular determinants of acute inflammatory responses to biomaterials. J Clin Invest 1996; 97: 1329-1334. 195. McNally AK, Anderson JM. Complement C3 participation in monocyte adhesion to different surfaces. PNAS 1994; 91:10119-10123. 196. Bernatchez SF, Parks PJ, Gibbons DF. Interaction of macrophages with fibrous materials in vitro. Biomaterials 1996; 17:2077-2086. 197. Kuwabara K, Ogawa S, Matsumoto M, Koga S, Clauss M, Pinsky DJ, Lyn P, Leavy J, Witte L, Joseph-Silverstein J, Furies MB, Torcia G, Cozzolino F, Kamada T, Stern DM. Hypoxia-mediated inducation of acidic/basic fibroblast growth factor and platelet-derived growth factor in mononuclear phagocytes stimulates growth of hypoxic endothelial cells. Proc Nat Acad Sci 1995; 92:4606-4610. 198. Van Otteren GM, Standiford TJ, Kunkel SL, Danforth JM, Strieter RM. Alterations of ambient oxygen tension modulate the expression of tumour necrosis factor and macrophage inflammatory protein-1 from murine alveolar macrophages. Amer J Respir Cell Mol Biol 1995; 13(4):399-409. 199. Leibovich SJ, Polverini PJ, Shepard HM, Wiseman DM, Shively V, Nuseir N. Macrophage-induced angiogenesis is mediated by tumour necrosis factor- α . Nature 1987; 329:630-632. 200. Patterson C, Perrella MA, Endege WO, Yoshizumi M, Lee M, Haber E. Downregulation of vascular endothelial growth factor receptors by tumour necrosis factor-α in cultured human vascular cells. Journal of Clinical Investigation 1996; 98(2):490-496. 201. Spyridopoulos I, Brogi E, Kearney M, Sullivan AB, Cetrulo C, Isner JM, Losordo DW. Vascular enothelial growth factor inhibits endothelial cell apoptosis induced by tumour necrosis factor-α: Balance between growth and death signals. J Mol Cell Cardiol 1997; 29:1321-1330. 202. Defife KM, Jenney CR, Colton E, Anderson JM. Confocal and light microscopic evaluation of silane surface-dependent macrophage development and IL-4 induced foreign body giant cell formation. Fifth World Biomaterials Congress 1996. 203. Kao WJ, Zhao QH, Hiltner A, Anderson JM. Theoretical analysis of in vivo macrophage adhesion and FBGC formation on PDMS, low density polyethylene and PEUs. J Biomat Res 1994; 28:73-79.

44 204. Kao WJ, Hiltner A, Anderson JM, Lodoen GA. Theoretical analysis of in vivo macrophage adhesion and FBGC formation on strained poly(etherurethane urea) elastomers. J Biomat Res 1994; 28:819-829. 205. Mathur A, Collier TO, Kao WJ, Wiggins M, Schubert MA, Hiltner A, Anderson JM. In vivo biocompatabiity and biostability of modified polyurethanes. J Biomater Res 1997; 36:246-257. 206. McNally AK, Anderson JM. The lymphokine IL-4 induces very large FBGC and syncytia from human macs in a material surface property dependent manner in vitro. Fifth World Biomaterials Congress 1996. 207. McNally AK, DeFife KM, Anderson JM. IL-4 induced macrophage fusion is prevented by inhibitors of mannose receptor activity. Am J of Pathol 1997; 149:975. 208. McNally AK, Anderson JM. IL-4 induces foreign body giant cells from human monocytes/macrophages. Amer J Pathol 1995; 47:1487-1499. 209. Essner R, Rhoades K, McBride WH, Morton DL, Economou JS. IL-4 downregulates IL-1 and TNF gene expression in human monocytes. J Immunol 1989; 142:3857-3861. 210. Lehn M, Weiser WY, Engelhor S, Gillis S, Remold HG. IL-4 inhibits H2O2 production and antileishmanial capacity of human culture monocytes mediated by fn-γ . J Immunol 1989; 143:3020-3024. 211. Lacraz S, Nicod L, Galve-de Rochemonteix B, Bauberger C, Dayer J-M, Welgus HG. Supression of metalloproteinase biosynthesis in human macrophages by interleukin-4. J Clin Invest 1992; 90:382-388. 212. Zhao Q, Topham N, Anderson JM, Hiltner A, Lodoen G, Payet CR. Foreign-body giant cells and polyurethane biostability: In vivo correlation of cell adhesion and surface cracking. J Biomat Res 1991; 35:177-183. 213. Anderson J. Inflammatory reaction: The nemesis of implants. In: Zilla P and Greisler H, eds. Tissue Engineering of Prosthetic Vascular Grafts. Austin: Landes Bioscience, 1998. 214. Forsyth EA, Hamdy MA, Neville RF, Sidawy AN. Proliferation and extracellular matrix production by human infragenicular smooth muscle cells in response to interleukin-1β. J Vasc Surg 1997; 26:1002-1008. 215. Makita S, Nakamura M, Yoshida H, Hiramori K. Autocrine growth inhibition of IL-1 beta-treated cultured human aortic smooth muscle cells: Possible role of nitric oxide. Heart Vessels 1996; 11:223-8. 216. Meredith JE Jr, Schwartz MA. Integrins, adhesion and apoptosis. Trends in Cell Biol 1997; 7:146-150. 217. Davis GE, Camarillo CW. Regulation of endothelial cell morphogenesis by integrins, mechanical forces and matrix guidance pathways. Exper Cell Res 1995; 216:113-123.

Tissue Engineering of Prosthetic Vascular Grafts 218. Vernon RB, Sage EH. Between molecules and morphology: Extracellular matrix and creation of vascular form. Amer J Pathol 1995; 147(4):873-883. 219. Polverini PJ. Cellular adhesion molecules: Newly identified mediators of angiogenesis. Amer J Pathol 1996; 148(4):1023-1029. 220. Iijima K, Yoshikawa N, Connolly DT, Nakamura H. Human mesangial cells and peripheral blood mononuclear cells produce vascular permeability factor. Kidney International 1993; 44:959-966. 221. Polverini PJ, Contran RS, Gimbrone MA Jr, Unanue ER. Activated macrophages induce vascular proliferation. Nature 1977; 269:804-806. 222. Anderson JM. Mechanisms of inflammation and infection with implanted devices. Cardiovasc Pathol 1993; 2:33S-41S. 223. Pepper MS, Belin D, Montesano L, Orci L and Vassalli JD. Transforming growth factor-b1 modulates basic fibroblast growth factor-induced proteolytic and angiogenic properties of endothelial cells in vitro. J Cell Biol 1990; 111:743-755. 224. van der Lei B, Dijk F, Bartels H, Jongebloed WL, Robinson PH. Healing of microvenous PTFE prostheses implanted into the rat femoral vein. Brit J Plastic Surg 1993; 46:110-115. 225. Sato Y, Rifkin DB. Inhibition of endothelial cell movement by pericytes and smooth muscle cells: Activation of a latent transforming growth factor—β1-like molecule by plasmin during coculture. J Cell Biol 1989; 109:309-315. 226. Clowes AW. Platelet-derived growth factor activity and mRNA expression in healing vascular grafts in baboons. J Clin Invest 1991; 87:406-414. 227. Jiang B, Yamamura S, Nelson PR, Mureebe L, Kent KC. Differential effects of platelet-derived growth factor isotypes on human smooth muscle cell proliferation and migration are mediated by distinct signaling pathways. Surgery 1996; 120:427-432. 228. Koyama N, Hart CE, Clowes AW. Different functions of the platelet-derived growth factor-α and—β receptors for the migration and proliferation of cultured baboon smooth muscle cells. Circ Res 1994; 75:682-691. 229. Sterpetti AV, Cucina A, D’Angelo LS, Cardillo B, Carvallaro A. Shear stress modulates the proliferation rate, protein synthesis, and mitogenic activity of arterial smooth muscle cells. Surgery 1993; 113:691-699. 230. Watkins MT, Sharefkin JB, Zajtchuk R, Maciag TM, D’Amore PA, Ryan US, Van Wart H, Rich NM. Adult human saphenous vein endothelial cells: Assessment of their reproductive capacity for use in endothelial seedingof vascular prostheses. J Surg Res 1984; 36:588-96.

CHAPTER 2 Noncompliance: The Silent Acceptance of a Villain Alexander M. Seifalian, Alberto Giudiceandrea, Thomas Schmitz-Rixen, George Hamilton

Introduction

A

rterial occlusive disease is a problem of epidemic proportions in our aging society with increasing need for vascular reconstructive surgery.1 The best results are achieved with autologous vein, but because this is often not available, historically there has been extensive research into production of a suitable substitute graft material.2 At present there are a handful of such vascular grafts that are either commercially available or under development, but unfortunately no graft material has yet performed satisfactorily.3,4 Baird and Abbott’s hypothesis of 1976 that a difference in circumferential compliance of a vascular graft and the host artery is detrimental to graft performance was eventually experimentally verified in 1987.5,6 The nature of compliance mismatch is complex, as it is determined by the compliance differences of the host artery, the anastomosis, and the graft itself. The hemodynamic consequences of mismatch include increased impedance and decreased distal perfusion as well as disturbed flow, turbulence and low shear stress rates.5 These changes could lead to the development of myointimal hyperplasia around the anastomosis, finally resulting in graft failure, particularly in small diameter vessels.7,8 Some efforts have been made over the years to engineer prosthetic grafts with compliance similar to that found in human artery.9,10 To date these attempts have failed due to a variety of reasons, explaining our current acceptance of highly noncompliant and inferior graft materials. This chapter will give an overview of the history of compliance in vascular grafting, a discussion of the nature of compliance and its measurement as a means of understanding the most recent attempts to engineer a more compliant graft.

Historical Overview Cardiovascular physiology started in 1628 with Harvey’s “De motu cordis et Sanguinis in Animalibus” in which he describes the circulation.11 Fifty years later in 1676 Robert Hook described the proportionality between stress and strain (“Ut Tensio Sic Vis” or, as the extension so the force) and the concept of elasticity of materials. Isaac Newton and his “Principia Matematica” was the first to relate elasticity (hence compliance) to wave velocity. His formula (equation 1) showed that velocity in a medium c is equal to the product of elasticity K and the density of the flowing medium ρ. (1) Thomas Young (1773-1829) investigated this principle in detail, both theoretically and experimentally, and gave his name to the elastic modulus. The Reverend Stephen Halea Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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in 1733 analyzed the role of a compliant vessel in maintaining a continuous blood flow throughout the pulsation of a cardiac cycle.12 The German translation of his work “Hemostatic”, introduced the important analogy in understanding the physiological role of distensibility in arteries, namely the Windkessel. This is an air reservoir which was fitted to fire engines of the 18th century in order to smoothen out the oscillations in flow of water due to intermittent pumping, thus ensuring a constant flow to quench the fire. Otto Frank in 1899 introduced a theoretical approach to explain the Windkessel effect on pressure waves.13 But it was not until 1950 through the work of Wormersley and McDonald that a rational mathematical approach to pulse flow and change in diameter over the cardiac cycle was reached which allowed comprehensive understanding of the hemodynamics of arterial flow.14,15 Parallel to these physiological understandings, vascular surgery evolved. The first great step which led to the foundation of vascular surgery was the pioneering work of Alexis Carrel in developing a technique for suturing blood vessels.16 This work published in 1902 lead to his Nobel Laureate in 1912. The first use of a graft to replace an excised segment of artery was reported by Goanes in Spain in 1906 when he bridged the defect with the popliteal vein of the patient.17 Later, homologous grafts were used but these attempts were frustrated by early graft degeneration and infections. In 1952 appeared the first report of a synthetic graft made from Vinyon-N, a new synthetic cloth used to manufacture ladies undergarments.18 In 1954 Vinyon-N was formally introduced to clinical practice. Since then many different materials have been introduced, with only Darcon and Teflon performing acceptably, such that virtually all modern prosthetic grafts are made from these materials.19 It was not until 1976 that Baird and Abbott postulated compliance mismatch as a major factor affecting graft patency.5,6 Since then a considerable amount of research has elucidated the nature of compliance mismatch, but as yet has not led to production of a more compliant graft.9,10,20

Physical Properties of the Vessel Wall Knowledge of the elastic properties of the arterial wall is essential to study the dynamics of the arterial system. During systole, with an increase in pressure the arterial wall circumference increases, with return during diastole to its previous dimension. This function of the vascular wall is of crucial physiological importance and due to a combination of elastic and viscous components inherent to arterial tissue. Solid materials have elasticity to varying extents as an inherent mechanical feature and this is defined as the ratio of applied stress to resultant strain. Young’s modulus expresses this quality of elasticity and is measured in units of dyne/cm2. Solid materials which regain their original dimensions when the stress is withdrawn are perfectly elastic, while those which retain the entire deformation are plastic. The property of viscosity distinguishes a fluid from a solid. When stress is applied to a fluid it will undergo viscous flow. The vessel wall exhibits properties of both an elastic solid and a viscous fluid and thus can be most properly described as being visco-elastic.

Tissue Engineering of Prosthetic Vascular Grafts

The deformation undergone by this class of material depends on the magnitude of the stress and on the rate at which it is applied. Vessel compliance, the reciprocal value of Young’s elastic modulus, is generally defined as the ratio of change in diameter over change in blood pressure (%/mm Hg x 10-2).20 It is defined by equation (2) where D and P are diameter and pressure, and the d and s subscripts denote diastole and systole respectively. (2) For convenience in comparing different vessels a standard reporting mean pressure of 100 mm Hg has been adopted. When thickness is a major fraction of the diameter of the vessel the distinction between inner and outer diameter becomes important. It must, therefore, be specified which diameter is used in calculating compliance. To compute the absolute compliance the pressure should be recorded at the exact point of diameter measurement, but this is not always the case, especially when calculating compliance in a clinical setting.21,22

Assessment of Compliance There are several strategies used to assess the elastic properties of vascular tissues. Classification of these methods is to some extent arbitrary and these so-called different methods share certain common principles. Two distinct methods are commonly used, namely longitudinal compliance, which assesses elasticity of a selected length of the vascular system, and circumferential compliance. This latter method is the most frequently used and assesses the compliance of the cross sectional area of interest of the vessel or graft. Longitudinal Compliance This technique is based on monitoring of pulse wave velocity (PWV) of blood down a given arterial pathway.23,24 The elastic component of the wall plays an important role in determining the velocity of propagation of a pulse wave. The Moert-Kortensweg equation (3) describes this relationship. In this equation E is Elastic module, h is wall thickness, R is radius and ρ is density of blood. (3) The compliance C can be computed from the above principle in equation 4. (4) This approach is based on the assumption that (1/2R) is small and that the vessel or grafts are filled with viscous liquid. Experimental evaluation of this methodology reveals an error of 16-24% in computation of compliance, thus requiring a correction factor.25 Since the pulse pressure and flow pulse propagates along a vessel with the same velocity, arterial compliance can therefore be indirectly measured by observing the flow pulse with Doppler ultrasound. Values for the time taken for this pulse to travel a unit length of the arterial pathway,

Noncompliance: The Silent Acceptance of a Villain

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in the absence of reflection from the periphery, can then be applied to equation 4 to give the average arterial compliance of a section of artery over which the pulse was measured. This technique requires a simultaneous use of two Doppler devices.26 To further complicate this method, increased PWV due to generalized stiffening of an artery or local plaque formation cannot be distinguished and accounted for. As a result, this approach has now been overtaken by recent noninvasive accurate assessments of crosssectional compliance. Circumferential Compliance The most popular and acceptable method of computing compliance is from measurement of changes in diameter over a cardiac cycle. The first approach for assessment of elasticity, and so compliance, is based on measurement of pressure and diameter curves over the cardiac cycle. In quasistatic compliance slow inflation and deflation of the blood vessels generate a pressure diameter curve. The arterial wall, being an anisotropic material, has a nonlinear elastic behavior. The incremental elastic modulus is calculated using equation 5. (5) Einc is the incremental elastic module, R is radius, o and i subscripts denote inner and outer radius respectively, σ is Poisson ratio, ∆R and ∆P are change in radius and pressure respectively. The second approach is based on assessment of excursions of diameter and pressure during a cardiac cycle and is known as dynamic compliance. Since strain relates not only to the magnitude of stress but also to the rate at which it is applied, it is logical to look at the elastic property in the physiological setting of pulsatile flow. A potential problem in simultaneous recordings of changes in diameter and pressure is the phase lag between the pressure and diameter curves, which is caused by the viscous component of the vascular wall. A complex elastic modulus has therefore been defined to formulate this concept. (6) Where nent and

is the elastic compois the viscous component. lm is

the average circumference and qm the average cross section area, P is the amplitude of the applied stress and Dl/l is the strain, w is the angular frequency, m is viscosity and f phase lag between the pressure and diameter curves. Changes in diameter and pressure over cardiac cycle can be measured at selected points on the vessel of interest, giving a more complete picture of the viscoelastic properties of the arterial wall. Recording of the pressure in vivo, however, requires insertion of an arterial catheter with its attendant possible complications. In the clinical setting, quasistatic compliance is more widely used. This technique

requires recording of systolic and diastolic pressures as well as noninvasively recording diameter over cardiac cycle. A wide spectrum of techniques have been used to measure the circumferential compliance in vitro as well as in the clinical setting, although less successfully. These methodologies include electrical, optical, ultrasound, magnetic resonance imaging and digital angiography. Assessment of Compliance Using Electrical Techniques Measurement of changes in diameter by placement of electrical resistance probes on blood vessels was first reported in 1960.27 The principle of this technique is based on placement of transducers on the arterial wall with continuous changes in diameter being electrically produced. The device which utilizes a miniature differential transformer measures changes in external diameter of blood vessels and was used in particular to map the mechanical properties at several sites along the aorta and some of its major branches.28 This methodology is highly invasive and effectively could only be used intraoperatively. In addition, the use of the device imposes mechanical restraint on the vessel because of the need to apply calipers or coils. Since a typical change in diameter of a blood vessel is in the order of about 5%, the degree of distortion and frictional force caused by these instruments led to acceptably high errors in measurements on small blood vessels. Furthermore these instruments measure external diameter only, giving no information about internal diameter movement, and all of this is can further be compounded by problems in identifying exactly the limit of the external wall in a native artery. Optical Techniques The main principle of this technique is application of light to measure diameter of the vessel. The vessel of interest is placed in a laser scanning system; this consists of a laser light and photocell detectors. In operation the laser tube emits a light source which after passing through appropriate objects sweeps along a particular axis at a known velocity. The presence of an artery in the beam path prevents light from reaching a photocell during a particular time. Knowing this time and sweep velocity calculations are made to compute the outside diameter of then artery in continuous mode.29 This system has been applied to assess the compliance of canine carotid artery, femoral artery and prosthetic grafts in vitro by placing the vessels in pulsatile flow circuits.30 The technique has reported accuracy of detection in the order of 13 µm changes in outer diameter of the artery.29 The disadvantage of this system is that it is restricted to experimental use only. In addition, it detects only changes in outer diameter and gives no information regarding inner wall movement, which may be different, particularly in thickened vessels. Ultrasonic Techniques Ultrasound has been the favored method of compliance measurement over the past two decades. There are two ways in which ultrasound has been applied to assess the compliance of blood vessels. The first measures the PWV along a segment while the other is based on the local distention

48

wave form of a local artery (local compliance).20 This latter method requires assessment of diameter of the artery under investigation at the onset of a cardiac cycle, and thus distention during the cardiac cycle under local pulse pressure. A radio frequency data acquisition system linked to a computer is required and M-mode echo images of the blood vessel will be displayed. The anterior and posterior walls are identified manually or automatically. The radio frequency signals from the ultrasound M-mode output over a cardiac cycle are digitized and relayed to a wall tracking system31 (see Fig. 2.1). From these signals end-diastolic and end-systolic intraluminal diameters can be easily determined and thus the maximum changes in diameter or distention for each beat. Blood pressure can be detected noninvasively and compliance computed. Ultrasound has inherent inaccuracies relating mainly to problems in recognizing the exact inner or outer diameter of the blood vessel.32,33 In vivo studies of the inter- and intraobserver variability reveal errors of 5% in measuring static diameter, and 10-15% in measuring pulsatile diameter changes.34 Increased angle of insonation by the transducer in the longitudinal axis will falsely increase diameter and its changes. Due to the tortuous nature of some vessels in vivo it is difficult to be exactly perpendicular to the vessel axis. The variability introduced by this problem in measuring diameter can affect the quality of results especially in looking at small caliber vessels and consequently small changes over the cardiac cycle. This uncertainly is partially overcome by the use of intravascular ultrasound, which has the major drawback, however, of being invasive.35 Magnetic Resonance Imaging Techniques Magnetic resonance imaging (MRI) is a noninvasive imaging modality that is rapidly gaining clinical acceptance, although widespread introduction has been delayed by its expense. MRI has been used directly to measure regional aortic compliance as well as total cardiac aortic compliance.36

Fig. 2.1. Displacement curves of the anterior (ant) and posterior (pos) walls of the common carotid artery in healthy young male (age 25 years). The bottom trace is the difference between the displacements of both arterial walls and represents the change in arterial diameter during the cardiac cycle. The first sign on the distention tracing refers to the trigger of the R-wave of the ECG (I), after which detection of end diastole and peak systole (I) starts. Abbreviation: b, heart beat; dist, distention; cca, common carotid artery. The data is obtained using a wall tracking system (Pie Medical, Masstricht, The Netherlands).

Tissue Engineering of Prosthetic Vascular Grafts

This technique is applicable clinically as well as experimentally for assessment of compliance. The disadvantage of MRI is that in vascular diseases such as atherosclerosis, focal lesions are formed; thus the compliance at one site may be very different from that at another. Further problems may occur in follow up studies to monitor the course of the disease or response to therapy, since relocation of the original site of measurement is very difficult. MRI, therefore, while useful in demonstrating accurately pathological disease, has limitations in these measurements. Digital X-ray Techniques Angiography, or radiographic imaging of blood vessels, has a well established role as the gold standard in the diagnosis of vascular disease and is widely used in clinical practice for obtaining high quality vessel images.37,38 Digital X-ray angiography has the necessary temporal and spatial resolution for volume blood flow measurement and estimation of diameter over cardiac cycles. Diameter estimates are either based on accurate vessel edge location or by densitometry, i.e., by integrating image brightness perpendicular to the axis of the vessel.39-41 The latter technique and principle yield a number proportional to the vessel’s area independent of lumen shape, in both healthy and diseased vessels. This technique has been used for estimation of compliance as well as blood flow in the vessel of interest.42,43 Measurement of vascular diameter and volume blood flow in vessels which are tortuous and which do not lie parallel to the imaging plane relies on accurate computation of the 3-dimensional (3D) path length of the vessel, X-ray magnification and the angle between the vessel axis and the X-ray beam. These factors can be computed from accurate 3D reconstruction of vascular geometry.42,44,45 The authors and collaborators have described a novel technique of measurement of vessel cross sectional area, and hence diameter, using densitometric methods applied to standard intraarterial digital subtraction angiograms.41,43 The technique is based

Noncompliance: The Silent Acceptance of a Villain

49

on image densitometry in which the integral intensity of the contrast bolus is computed along a profile perpendicular to the projection of the vessel axis. As illustrated in Figure 2.2, the true cross-section A is related to the densitometric measure a by:

ization, X-ray angiography is still the modality of choice for critical morphological vascular studies. That X-ray angiography has not been widely used for measuring blood flow is due in part, we believe, to the use of inappropriate algorithms for processing the imaging digital data.46,47 The method does, however, have great potential, especially when combined with lower dose digital subtraction angiography and the new nonionic contrast agents.48 Mini-puncture needles and catheters have led to increased safety of the technique and the equipment and expertise is available in most centres.49

(7) The X-ray magnification factor M and angle θ between the vessel axis and the X-ray axis are computed from the 3D reconstruction of the vascular configuration from two views. The densitometric calibration constant K relates the image gray value to the mass of iodine integrated along the X-ray path from X-ray focus to image. These were obtained from data generated from 3D reconstructions of biplanar X-ray angiographic data.42 Although it requires vascular catheter-

Synthetic and Biological Grafts Prosthetic grafts fare well when used in the aortic or aortofemoral position, but are much less successful for bypasses below the inguinal ligament, with the patient’s own Fig. 2.2. The true cross-section is related to the integral of image intensity along a profile perpendicular to the projection of the vessel axis. See text for details.

Table 2.1. Summary of results from literature for five years cumulative patency rates of infrainguinal reconstruction with autogenous vein study Research Workers

Taylor et al 199051 Berganini et al 199152 Donaldson et al 199253

No of limbs

Operative mortality

Primary graft patency (%)

Secondary graft patency (%)

Limb salvage (%)

Patient survival (%)

516 361 440

1 3 2

75 63 72

80 81 83

90 86 84

28 57 66

50

Tissue Engineering of Prosthetic Vascular Grafts

saphenous vein giving the best long term patency rates.2,50 Because of dismal patency rates, prosthetics are not used in coronary bypass grafting. Primary 5 year patency rates for autogenous vein bypasses range between 63-75%, and secondary patency rates between 80-83%, leading to limb salvage rates which range from 84-92% (Table 2.1).51-53 These are good clinical results but, unfortunately, in up to 40% of patients autogenous vein may not be available or be inadequate because of coexisting disease such as varicose veins or previous venous thrombosis. In these circumstances a prosthetic graft is the only possible option.3 Since 1950, prosthetic grafts have been used in a variety of locations for arterial reconstruction. Knitted and woven Dacron grafts provided favorable results in high flow large caliber, vessels but performed poorly in more challenging small caliber low flow conditions characteristic of infrainguinal bypasses, particularly to the below knee arteries. To date the most successful prosthetic grafts in this situation have been expanded polytetrafluoroethlylene (PTFE) and the Dardik biograft or human umbilical vein tanned with glutaraldehyde.54,55 The human umbilical vein graft was introduced at roughly the same time as PTFE but never gained broad acceptance, primarily because of a propensity for dilatation and aneurysm formation as early as 2 years after implantation, occurring in up to 57% of the grafts.54,56,57 This phenomenon of graft dilatation is common to all biografts and is attributed to a combination of deterioration of collagen crosslinks with time, and proteolytic digestion by host enzymes almost certainly mediated by an immune response.58 In vitro testing of 2 types of biografts, a bovine carotid and the human umbilical vein graft, with long term perfusion with proteolytic enzymes, demonstrated changes in the structural integrity of the graft. This was detected by a degree in compliance, thus providing experimental data for the role of proteolytic digestion in these grafts failures.59 Despite this, clinical experience with the human umbilical vein gives a 5 year patency with approximately 60% for femoral popliteal grafting with up to 80% in the above knee location.60,61 Thus these grafts perform clinically well in the short term and have been recommended for use in patients with a short life expectancy. Results achieved with PTFE bypass grafting are less satisfactory particularly so onto the below knee tibial vessels. On average the long term patency rates (> 3 years) in

Causes of Graft Failure Causes of graft failure are complex and include technical failure, poor selection, compliance mismatch, primary thrombotic failure related to low flow environments, development of intimal hyperplasia, late failure due to graft gen-

Table 2.2. Summary of results from literature relationship between compliance vs. patency rate.51,54,63 Graft type

Host artery Saphenous vein Umbilical vein Bovine heterograft Dacron PTFE

Compliance (%mm Hg x 10-2) 5.9 ± 4.4 ± 3.7 ± 2.6 ± 1.9 ± 1.6 ±

0.5 0.8 0.5 0.3 0.3 0.2

Patency %

75 60 59 50 40

6 Compliance (%/mm Hg x 10-2)

Fig. 2.3. Data reported from compliance of the various biological and prosthetic grafts vs. patency rate.

infrapopliteal reconstructions is dismal.62,63 PTFE grafting into the popliteal artery at the knee shows a slightly better 5 year patency rate of 40%.61-63 Recently, better patency rates, even to the tibial level, have been reported with the use of PTFE bypass grafting onto an interposition vein collar or patch which is placed onto the distal artery. In part this may be due to improvement of the compliance match between the host artery and the graft.64 Compliance mismatch has an important role in graft failure. Various research workers have shown that compliance of biological conduit is significantly greater than that of prosthetic material; this compares well with patency rates for the two different graft materials (Table 2.2).51,54,63 It appears that as compliance mismatch increases, patency decreases; linear regression analysis of this data shows a highly significant correlation for this association, as shown in Figure 2.3. The patency data for the different grafting materials used in the femoro-popliteal position used in this study were obtained from several recent large studies and compliance data from a study performed by Abbot et al in Boston.51,54,55,63 Finally, prosthetic grafts have poor dynamic compliance profiles in comparison to arteries with no increase in compliance occurring at low pressure, an important physiological response to shock (Fig 2.4).20

5

Y = 0.09 x -2.0 r2 = 0.877, p = 0.019

■ ● ◆ ▼ ▲

4 3 2 1 30

40

50 60 Patency (%)

70

80

Umbilical vein Bovine heterograft Dacron PTFE Saphenous vein

Noncompliance: The Silent Acceptance of a Villain

51

Fig. 2.4. Compliance-pressure curve comparing human superficial femoral artery (●) with PTFE 6 mm graft (❏). Note that an important feature of human vessel is the increased compliance found at lower physiological pressures. This has the important benefit of preserving pulsatile energy in situations of shock, thus optimizing flow. This dynamic change in compliance is totally absent in PTFE and Dacron grafts, and may be a further reason for prosthetic graft thrombosis occurring during hypotensive states.

8

Compliance

6

4

2

0 40

60

80

100

120

Mean pressure (mm Hg)

eration and late failure due to progression of native vessel arterial sclerosis. More than one of these factors may be implicated in graft failure. Much research has been focused on these various factors, with particular emphasis on reducing the thrombogenicity of grafts either by modifying the luminal surface or by endothelial cell seeding. Much less effort has been directed into the role of compliance in graft failure. Possibly, this is because of limitations in materials and material technology. There is much clinical data which emphasizes the importance of compliance mismatch and certainly it is established experimentally that current noncompliant grafts owe their inferior performance to lack of distensibility when compared to native arteries.6 It is most important to consider compliance mismatch as composed of two major parts, namely tubular and anastomotic.

Tubular Compliance Tubular compliance mismatch refers to the imparity of elasticity between the prosthetic conduit and the native artery. Essentially this results in an adverse hemodynamic effect and reduced distal perfusion, as was previously discussed. A compliant wall acts as an elastic reservoir absorbing energy during systole, which is released during diastole. A rigid vessel wall consequently diminishes the pulsatile component of the diastolic recoil, thus reducing the energy available for distal perfusion. Impedance is the term given to resistance to pulsatile flow. Change in impedance occurs at the interface between a compliant artery and noncompliant graft, and this results in propagation of less than 60% of the pulsatile energy.65 Optimal organ perfusion depends on pulsatile flow, and it has been shown that change from pulsatile to steady state flow causes peripheral resistance to increase by 10%.66 Wave reflection of the graft artery interface also leads to increased

velocity gradients and turbulence. As a consequence of these mechanical effects, vibratory weakening of the arterial wall has been postulated as a cause of endothelial damage, intimal hyperplasia and even anastomotic aneurysm formation.

Anastomotic Compliance Mismatch Anastomotic compliance mismatch refers to the variation in compliance occurring specifically at the anastomotic site. The change in compliance from native artery and prosthetic material is not monotonic because the creation of an anastomosis always generates a focal decrease in diameter and a drop in compliance. This is partially determined by the elasticity, or rather rigidity, of the suture material, and partly by the surgical technique, depending on whether a continuous or interrupted anastomosis is used. Interrupted anastomotic suture techniques give more compliant anastomoses. There is also a paradoxical increase in compliance of about 50% which occurs within a few millimeters on either side of the suture line (Fig. 2.5). This characteristic is known as the para-anastomotic hypercompliance zone (PHZ).20,67,68 It has been hypothesized that PHZ may be responsible for the intimal hyperplasia which characteristically develops at the same area as this hyper-compliance. Thus mismatch of elastic properties around the anastomosis may act to promote intimal hyperplasia in at least 3 different ways: 1. Compliance mismatch between the artery and graft may lead to a region of excessive mechanical stress, possibly resulting in wall injury, a major event in the initiation of intimal hyperplasia;6,69,70 2. Cyclic stretching has a positive influence on replication of vascular smooth muscle cells and production of extracellular matrix.70,71 Experimentally, it has been demonstrated that vascular smooth muscle cell

52

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 2.5. Para-anastomotic hypercompliant zone (PHZ). Typical compliance profile of a continuous anastomosis: Compliance vs. distance. Anastomosis is at 0 mm.

start to produce extracellular matrix and to replicate when they are subjected to high levels of distension.72 Also, animal studies have confirmed that intimal thickening is associated with increased tangential stress;73 3. Flow studies show that a sudden increase in compliance is associated with enhanced particle residence time, flow separation and stasis leading to low shear stress conditions.74,75 Also, from animal studies it is known that high levels of shear stress have a deleterious effect on endothelial cells but, most importantly, there is also mounting evidence that intimal thickening occurs in areas of low shear rate. An example of this phenomenon in humans is of the sudden increase in compliance found in the carotid bulb where the above hemodynamic changes occur. Since this is a site with predilection for the occurrence of atherosclerosis, one could postulate that increased compliance with prolonged particle residence time, flow separation and low shear stress conditions can cause damage, allowing initiation of arteriosclerosis in man.

Signal Transduction Pathways and Flow Flow is an important modulator of vascular structure and function, probably mediated through the effects of the function on the endothelial cell.76,77 Pathways by which flow acts on the vessel wall are poorly defined, but an essential role is attributed to the endothelial cell.78-80 Nitric oxide (NO) plays an important role in this pathway, and it has been shown that specific potassium channel antagonists can block flow mediated vasodilatation and nitric oxide release, suggesting that the signal transduction pathway underlying flow-mediated vasodilatation may be the activation of the potassium channel on the endothelial cell membrane.73 The

finding of free calcium concentrations proportional to the shear rate suggests an important role of calcium in the signal transduction pathway.81 Shear rate dependent release of growth factor is reflected by a varying expression of multiple growth factor-related genes when wall shear rates are changed.82 A possible mechanism underlying the increased expression of growth factor genes may be a recently described shear rate-induced transcription factor which may not only regulate the expression of platelet derived growth factor (PDGF) genes but also of tissue plasminogen activator (tPA), intracellular adhesion molecule 1 (ICAM-1) and transforming growth factor-b1 (TGF-β1).83 A shear stress dependent release of PDGE-β mRNA has also been observed in cultured human umbilical vein cells.84 These findings begin to illuminate the link between mechanical effects and changes and the development of intimal hyperplasia in human vessels and bypass grafts.

Development of a Graft with Better Compliance The ideal characteristics of a graft are biocompatability, a nonthrombogenic surface, physical durability, elastic or compliant properties similar to those of a native artery, resistance to infection and ease of implantation. Advances in manufacture of biomaterials and tissue engineering are leading to developments in the manufacture of more compliant grafts. What is not known, however, is the range of compliance required to give the ideal match between the graft and the host artery. Arteries change with age with loss of elastin and consequent reduction of elasticity and thus compliance. When the ravages of arteriosclerosis with medial calcification occurring as a result of age are added to these effects, the compliance of the host artery in a patient requiring a bypass may be dramatically reduced (Fig. 2.6).20 Thus the aim should be to develop a graft which is pulsatile,

Noncompliance: The Silent Acceptance of a Villain

53

Fig. 2.6. Tendency of correlation of compliance at mean pressure 100 mm Hg with increasing age on human pathology without calcification and major plaque formation.

6

Compliance

5

4

3

2 1

0 30

40

50

60

70

80

Age (years) Fig. 2.7. Compliance versus pressure excursion of different CronoFlex graft types (1-7).

but it may even be detrimental to have a graft which is too compliant. Study of typical compliances in aged and diseased human arteries, coronary, femoral, popliteal and tibial, is required before the ideal graft can be designed. Development of more compliant grafts is currently occurring, either using biomaterials or by tissue engineering of vascular grafts. An example of developments in biomaterials is the manufacture of grafts using polyurethane. Grafts have been made from this material for quite some time and the major problem has been biodegradability and subsequent aneurysm formation. A novel formulation of a polyurethane graft, namely CronoFlex (Poly Medica Industries, Tarvin, Cheshire, UK) has recently allowed produc-

tion of artificial blood vessels with specific compliance characteristics (Fig. 2.7).10 In this process the graft material is made using either a single shot method (Fig. 2.8)9 or a prepolymer technique (Fig. 2.9).9 In a single shot method, all reactants are added simultaneously and the final structure is thus dependent on the relative reactivity of the formulation. In the prepolymer method the ingredients are added sequentially, allowing close engineering and design of the final structure. Using a coagulation precipitation technique, porosity of the graft can be varied to form either micro- or macroporous structures. By varying the quantity and type of chemical groupings within the polyurethane polymer, it is possible to manipulate the level of crosslinking and thus

54

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 2.8. Production of polyurethane graft using single shot polymerization method. See text for details.

Fig. 2.9. Production of polyurethane graft using prepolymer polymerization method.

control mechanical properties. Thus a graft can be designed and modulated, varying not only the characteristics of the polymer but also the degree of “sponginess” of the graft wall by modulating the size of the pores or bubbles in the wall of the graft. This allows design of a graft to a specific compliance. This particular formulation of polyurethane grafts, known colloquially as the “bubble graft,” has demonstrably higher compliance characteristics than others (Fig. 2.10).10 By modulation of the outer wall of the graft using various strategies to strengthen it, kinking as the graft crosses the knee joint can be avoided. Extensive animal studies have shown no tendency for the CronoFlex graft to develop aneurysms.

A further important area in design of such a biomaterial is in selection of optimal luminal characteristics. It is possible to design the lumen of the graft so that it is inherently nonthrombogenic by incorporation, for example, of heparin or nonthrombogenic polymers. The alternative approach is to make the lumen more attractive for endothelial cell adhesion and seeding. A simple example of such a methodology is the CronoFlex graft lumen, which is constituted of many pits and valleys. Experimental evidence suggests that endothelial cells adhere better and are more resistant to the effects of flow shear stress because they can shelter within the pits, and proliferate upwards to endothelialize the graft (Fig. 2.11).85

Noncompliance: The Silent Acceptance of a Villain

55

Fig. 2.10. Scanning electrom microscopy cross-section of ChronoFlex graft. (A) Nonreinforced (original magnification x12.4).

Fig. 2.10.(B) Externally reinforced (original magnification x110.5).

Tissue engineering holds much promise for the development of more compliant grafts. The relative roles of collagen, elastin and smooth muscles in the mechanical characteristics of an artery are now well understood. Experimentally, the development of arterial grafts using combinations of smooth muscle cells, collagen and endothelial cells is possible. Major problems persist in terms of the source of these cells. Ideally these should be from the patient, but certainly in peripheral vascular grafting the time and logistics involved would mitigate against the widespread use of this new technology. The important role of the immune response in eventual biodegradability means that such tissue engineered grafts made from nonautologus cells would require the use of either immunosuppressive therapy, or cells of fetal ori-

gin, which are said to be nonimmunogenic. Little data is available as yet on the compliance characteristics of these materials. Certainly this approach would hold the greatest promise of developing a graft with the necessary visco-elastic properties comparable to those of human arteries.

The Future In the very near future it will be possible to move away from the use of the rigid materials, mostly Daron and PTFE, which are currently used in bypass grafting. Initially, prosthetic biomaterials with suitably designed and compatible compliance characteristics will be used, with clinical trials expected to begin within 2 years. The further addition of endothelial cell seeding to these more compliant grafts holds

56

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 2.11. Scanning electron microscopy (original magnification x660) of fibronection coated PTFE (A, top) and CPU (B, bottom) grafts after exposure to flow for 6 h.

the promise of providing an off the shelf vascular conduit with performance approximating that of the human saphenous vein. In the longer term, perhaps 5-10 years from now the problems of designing tissue engineered grafts, perhaps from the patient’s own cells, will have been conquered, with the exciting probability that cardiac and vascular surgeons will have available a conduit as good as a human artery to rescue failing hearts, organs and limbs. Acknowledgments We wish to thank Mr. Alan Edwards, Poly Medica Industries, Tarvin, Cheshire, UK for providing us with polyurethane graft figures.

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Noncompliance: The Silent Acceptance of a Villain 7. Kinley CE, Paasche PE, MacDonald AS. Stress at vascular anastomosis in relation to host artery: Synthetic graft diameter. Surgery 1974; 75(1):28-30. 8. Madras P, Ward C, Johnson W, Singh P. Anastomotic hyperplasia. Surgery 1981; 90(5):992-3. 9. Reed AM, Potter J, Szycher M. A solution grade biostable polyurethane elastomer: CronoFlex AR. Journal of Biomaterials Applications 1994; 8:210-36. 10. Edwards A, Carson RJ, Bowald S, Quist WC. Development of a microporous compliant small bore vascular graft. Journal of Biomaterials Applications 1995; 10:171-87. 11. Harvey W. (1628) De Motu Cordis, Frankfurt: William Fitzer. Translated as movement of the heart and blood in animals by Franklin, K.J. (1957). Oxford: Blackwell Scientific Publications. 12. Hales S. (1733) Statical essays: containing haemastaticks, reprinted 1964, No. 22, History of Medicine Series, Library of New York Academy of Medicine. New York: Hafner Publishing. 13. Frank O. Die Theorie der Pulswelle. Z Biol 1927; 85:91-130. 14. Womersley JR. Oscillatory flow in arteries: The constrained elastic tube as a model of arterial flow and pulse transmission. Phys Med Biol 1957; 2:313-23. 15. McDonald DA. The realtion of pulsatile pressure to flow in arteries. J Phsysiol 1955; 127:533-52. 16. Carrel A. La technique operatoire des anastomoses vasculaires et la transplantation des visceres. Lyon Med 1902; 98:859-64. 17. Goyanes J. Nuevos trabajos de cirugia vascular, substitucion plastica de las arterias por las venas or arterioplastia venosa, applicada, como nuevo metodo, al tratamiento de los aneurismas. Siglo Med 1906; 53:546-9. 18. Voorhees AB, Jaretzke AL, Blakemore AH. The use of tubes constructed from Vinyon-N cloth in briding arterial defects. Ann Surg 1952; 135:332-6. 19. Turner RJ, Sawyer PN. Modern Vascular Grafts. In: Vascular Graft Development: An Industrial Perspective. New York: McGraw-Hill,; 1985:75-103. 20. Schmitz-Rixen T, Hamilton G. Compliance: A critical parameter for maintenance of arterial reconstruction? In: Greenhalgh RM, Hollier LH, eds. The maintenance of arterial reconstruction. London: WB Saunders, 1991:23-43. 21. Murgo JP, Westerhof N, Giolma JP, Altobelli SA. Aortic input impedance in normal man: Relationship to pressure wave forms. Circulation 1980; 62:105-16. 22. O’Rourke MF. Systolic blood pressure: Arterial compliance and early wave reflection and their modification by antihypertensive therapy. J Hum Hypertens 1989; 3:47-52. 23. O’Rourke M. Arterial compliance and wave reflection. Arch Mal Coeur 1991; 84(III):45-8. 24. Lehmann ED. Elastic properties of the aorta. The Lancet 1993; 342:1417. 25. Bergel, DH. The visco-elastic properties of the arterial wall. Ph.D. Thesis, London University, 1960. 26. Bertram CD. Ultrasonic transit-time measurement for arterial diameter measurement. Med Biol Eng Comp 1977; 15:489-99. 27. Patel DJ, Mallos AJ, Fry DL. Aortic mechanics in the living dog. J App Physiology 1961; 16:283-99. 28. Peterson LH, Jenesen RE, Parnell J. Mechanical properties of arteries in vivo. Circulation Research 1960; 8:622-39.

57 29. Brant AM, Rodgers VGJ, Borovetz HS. Measurement in vitro of pulsatile arterial diameter using a helium-neon laser. J Appl Physiol 1987; 62:679-83. 30. Megermann JM, Hasson JE, Warnock D, L’Italien GJ, Abbott WM. Noninvasive measurments of nonlinear arterial elasticity. American Journal of Physiology 1986; 250:H181-8. 31. Hoeks APG, Brands PJ, Smeets FAM, Reneman RS. Assessment of the distensibility of superficial arteries. Ultrasound Med Biol 1990; 16(2):121-8. 32. Jaffe CC. Doppler applications and limit of the method. Clin Diagn Ultrasound 1984; 13:1-10. 33. Beach KW. The evaluation of velocity and frequency accuracy in ultrasound duplex scanners. Journal of Vascular Technology 1990; 14(5):214-20. 34. Hansen D, Bergqvist D, Mangell P et al. Non-invasive measurement of pulsatile vessel diameter change and elastic properties in human arteries—a methodological study. Clin Physiol 1993; 13:631-43. 35. Hansen ME, Yual EK, Megerman J. et al. In vivo determination of human arterial compliance: Preliminary investigation of a new technique. Cardiovascular & Interventional Radiology 1994; 17(1):22-6. 36. Mohiaddin RH, Longmore SR. MRI studies of atherosclerotic vascular disease: Structural evaluation and physiological measurements. Brit Med Bull 1989; 45:968-90. 37. Pond GD, Osborne RW, Capp MP et al. Digital subtraction angiography of peripheral vascular bypass procedures. AJR 1982; 2:279-81. 38. Guthaner DF, Wexler L, Enzmann DR et al. Evaluation of peripheral vascular disease using digital subtraction angiography. Radiology 1983; 147(2):393-8. 39. Brown BG, Bolson E, Bolson E, Frimer M, Dodge HT. Quantitative coronary arteriography. Estimation of dimensions hemodynamic resistance and atheroma mass of coronary artery lesions using the arteriogram and digital computation. Circulation 1977; 55:329-37. 40. Reiber JHC, Kooijman CJ, Slager CJ. Computer assisted analysis of the severity of obstructions from coronary cineangiograms: A methodological review. Automedica 1984; 5:219-38. 41. Hawkes DJ, Colchester ACF, de Belder MA et al. The measurement of absolute lumen cross sectional area and lumen geometry in quantitative angiography. In: ToddPokropek AE, Viergever MA, eds. Medical images: Formation, Handling and Evaluation. Nato ASI Series. Heidelburg: Springer-Verlag, 1992:609-26. 42. Hawkes DJ, Seifalian AM, Colchester ACF, Iqbal N, Hardingham CR, Bladin CF, Hobbs KEF. Validation of volume blood flow measurements using three-dimensional distance-concentration functions derived from digital xray angiograms. Investigative Radiology 1994; 29:434-42. 43. Seifalian AM, Hawkes DJ, Giudiceandra A. et al. A novel technique of blood flow and compliance measurement using digital subtraction angiography. In: Greenhalgh RM, ed. Vascular imaging for surgeons. London: W.B.Saunders Company Ltd, 1995:51-70. 44. Hawkes DJ, Hardingham CR, Seifalian AM et al. Recent advances in extracting quantitative information from xray angiographic data. Innov Tech Biol Med 1992; 13:99-107. 45. Hawkes DJ, Mol CB, Colchester ACF. The accurate 3-D reconstruction of the geometric configuration of vascular trees from x-ray recordings. In: Guzzardi R, ed. Physics

58 and engineering of medical imaging. NATO ASI. Holland: Martinus Nijhoff Publications, 1987:250-258. 46. Seifalian AM, Hawkes DJ, Colchester ACF, Hobbs KEF. A new algorithm for deriving pulsatile blood flow waveforms tested using simulated dynamic angiographic data. Neuroradiology 1989; 31:263-9. 47. Du Boulay GH, Brunt J, Colchester A et al. Volume flow measurement of pulsatile flow by digitised cine angiography. Acta Radiologica 1987;Suppl 13:59-62. 48. Bettman MA, Morris TW. Recent advances in contrast agents. RCNA 1986; 24(3):347-57. 49. Cope C. Minipuncture angiography. RCNA 1986; 24(3):359-67. 50. Crawford ES, Bomberger RA, Glaeser DH et al. Aortoiliac occlusive disease: Factors influencing survival and function following reconstructive operation 0ver a twenty-five year period. Surgery 1981; 90(6):1055-67. 51. Taylor LM, Edwards JM, Porter JM. Present status of of reversed vein bypass grafting: Five years result of a modern series. J Vasc Surg 1990; 11(2):193-205. 52. Bergamini TM, Towne JB, Bandyk DF. Experience with in situ saphenous vein bypasses during 1981 to 1989: Determinant factors of long term patency. J Vasc Surg 1991; 13(1):137-47. 53. Donaldson MC, Mannick JA, Whittemore AD. Causes of primary graft failure after in situ saphenous vein bypass grafting. J Vasc Surg 1992; 15(1):113-8. 54. Dardik H, Ibrahim IM, Dardik I. Evaluation of glutaraldehyde tanned human umbilical cord vein as a vascular prosthesisfor bypass to the popliteal, tibial and peroneal arteries. Surgery 1978; 83(5):577-88. 55. Walden R, L’Italien G, Megerman J. Matched elastic properties and successful arterial grafting. Arch Surg 1980; 115:1166-1169. 56. Dardik H, Ibrahim IM, Sussman B. Biodegradation and aneurysm formation in umbilical vein grafts: Observations and a realistic strategy. Ann Surg 1984; 199:61-8. 57. Hasson JE, Newton WD, Waltman MD et al. Mural degenaration in the glutaraldehyde-tanned umbilical vein graft: incidence and implications. J Vasc Surg 1986; 4:243-50. 58. Mattila SP, Fogarty TJ. Antigenicity of vascular heterograft. J Surg Res 1973; 15:81-90. 59. Hamilton G, Megerman J, L’Italien GJ et al. Prediction of aneurysm formation in vascular grafts of biologic origin. J Vasc Surg 1988; 7(3):400-8. 60. McCollum C, Kenchington G, Alexander C. PTFE or HUV for femoro-popliteal bypass: A mnulti-center trial. Eur J Vasc Surg 1991; 5(4):435-43. 61. Aalders GJ, van Vroonhoven TJM. PTFE versus HUV in above knee femoro-popliteal bypass. Six year results of a randomized clinical trial. J Vasc Surg 1992; 16:816-23. 62. Witthermore AD, Kent KC, Donaldson MC. What is the proper role of polytretrafluoroethlene grafts in infrainguinal reconstruction? J Vasc Surg 1998; 10(3):299-305. 63. Veith FJ, Gupte SK, Ascer E et al. Six-year prospective muticenter randomized comparison of autologous saphenous vein and expanded polytetrafluoroethylene graft in infrainguinal arterial reconstruction. J Vasc Surg 1986; 3(1):104-14. 64. Beard JD, Benveniste GL, Miller JH, Baird RN, Hoskins PR. Haemodynamics of the interposition vein cuff. Br J Surg 1986; 73:823-5. 65. Strandness DE, Summer DS. Hemodynamics for Surgeon. New York: Grune &Stratton, Inc. Pub., 1975.

Tissue Engineering of Prosthetic Vascular Grafts 66. Giron F, Birtwell WS, Soroff HS, Deterling RA. Hemodynamic effects of pulsatile and nopulsatile flow. Arch Surg 1966; 93(5):802-10. 67. Abbott WM, Megermann JM. Adaptives responses of arteries to grafting. Journal of Vascular Surgery 1989; 9:377-9. 68. Hasson JE, Megermann JM, Abbott WM. Increased compliance near vascular anastomosis. Journal of Vascular Surgery 1985; 2:419-23. 69. Bassiouny HS, White S, Glagov S et al. Anastomotic intimal hyperplasia: Mechanical injury or flow induced. Journal of Vascular Surgery 1992; 15:708-17. 70. Clowes AW, Clowes MM, Fingerle J, Reidy MA. Kinetics of cellular proliferation after arterial injury. Laboratory Investigation 1989; 60:360-4. 71. Sottiurai VS, Sue SL, Feinberg EL II et al. Distal anastomotic intimal hyperplasia: Biogenesis and etiology. Eur J Vasc Endovasc Surg 1988; 2:245-56. 72. Predel HG, Yang Z, von Segesser L et al. Implication of pulsatile stretch on growth of saphenous vein and mammary artery smooth muscle. Lancet 1992; 340:878-9. 73. Cooke JP, Rossitch E, Andon NA et al. Flow activates a endothelial potassium channel to relaese an endogenous nitrovasodilator. Journal of Clinical Investigation 1991; 88:1663-71. 74. Stewart S, Lyman DJ. Effects of vascular graft/natural artery compliance mismatch on pulsatile flow. J Biomech 1992; 25:297-310. 75. Ojha M, Cobbold RS, Johnston KW. Haemodynamics of a side to end proximal arterial anastomosis model. Journal of Vascular Surgery 1993; 17:646-55. 76. Kamiya A, Ando J, Shibata M, Masuda H. Roles of fluid shear stress in physiological regulation of vascular structure and function. Biorheology 1988; 25(1-2):271-8. 77. Hofstra L, Bergmans DC, Leunissen KM, Hoeks AP, Kitslaar PJ, Tordoir JH. Prosthetic arteriovenous fistulas and venous anastomotic stenosis: influence of a high flow velocity on the development of intimal hyperplasia. Blood Purif 1996; 14(5):345-9. 78. Luscher TF. Endothelial control of vascular tone and growth. Clin Exp Hypertens 1990; 12:897-902. 79. Luscher TF. Imbalance of endothelium-derived relaxing and contracting factors. A new concept in hypertension. Am J Hypertens 1990; 3:317-30. 80. Koo E, Gotlieb AI. Endothelial stimulation of intimal cell proliferation in a porcine aortic organ culture. American Journal of Pathology 1989; 134:497-503. 81. Shen J, Luscinskas FW, Conolly A et al. Fluid shear stress modulates cytosolic free calcium in vascular endothelial cells. American Journal of Physiology 1992; 31:c384-c390 82. Taubmann MB, Rollins BJ, Poon M et al. JE mRNA accunulates rapidly in aortic injury and in platlet-derived growth factor-stimulated vasular smooth muscle cells. Circulation Research 1992; 70:314-25. 83. Resnick N, Collins T, Aktinson W et al. Platlet-derived growth factor beta chain contains a cis-acting fluid shearstress-responsive element. Proc Natl Acad Sci USA 1993; 90:4591-5. 84. Hsieh HJ, Li NQ, Frangos JA. Shear stress increases endothelial-platelet-derived growth factor mRNA levels. American Journal of Physiology 1991; 260:H642-H646. 85. Giudiceandra A, Seifalian AM, Krijgsman B, Hamilton G. Effect of prolonged pulsatile shear stress in vitro on endothelial cell seeded PTFE and compliant polyurethane vascular grafts. Eur J Vasc Endovasc Surg 1998; 15:147-154.

PART II Biolized Prostheses: Surface Healing

CHAPTER 3 Endothelial Cell Seeding: A Review Steven P. Schmidt, Gary L. Bowlin

Introduction

T

he field of Tissue Engineering offers the promise of further elucidating and clarifying the interactions between endothelial cells and smooth muscle cells. This understanding is central to the creation of a successful small diameter vascular graft—a goal which has heretofore eluded researchers. Small diameter vascular grafts are defined as being less than 6 mm in internal diameter. These grafts fail in patients from thrombosis and intimal hyperplasia, processes which may, at least in part, originate from aberrant endothelial and smooth muscle cell interactions. Clarification of the biology of these two cell types and their interactions is critical in order to develop therapeutic solutions to the problems of small vessel prosthetic grafting. These solutions will also undoubtedly involve further elucidation of the roles of extracellular matrix in endothelial cell and smooth muscle cell biology, as well as of the responsiveness of these cells to a variety of cytokines. An endothelial cell lining does not naturally develop on the luminal surfaces of prosthetic vascular grafts in humans.1,2 This is in contrast to other animal species that have been studied in which neoendothelialization of graft luminal surfaces occurs by endothelial cell proliferation from perianastomotic artery, the microvessels of graft interstices, or circulating progenitor endothelial cells.3 Because endothelial cells release molecules that modulate coagulation, platelet aggregation, leukocyte adhesion and vascular tone, the absence of these cells lining prosthetic grafts, in combination with the inherent hostility of the biomaterial/ blood contacting surface, predisposes these grafts to platelet deposition leading to thrombosis. Although the precise mechanisms of the genesis of anastomotic intimal hyperplasia are still being defined, endothelial cell/smooth muscle cell dysfunctions are thought to be involved. Early investigators in the field of small-diameter graft development sought to promote graft endothelialization by transplantation of autologous endothelial cells onto vascular grafts prior to their implantation—a process that became known as endothelial cell seeding. The hypothesis underlying this research effort was really quite simple—that is, by promoting the establishment of the patient’s own endothelial cells on the blood-contacting surface of a vascular prosthesis, a “normal” endothelium-lined neointima would form on the graft whose biology would counteract rheologic, physiologic, and biomaterial forces promoting graft failure. In retrospect this simple hypothesis seems deceptively naive. A review of the history of endothelial cell seeding research has been published which details early animal and clinical studies.4 The results of these early animal studies were really quite promising. This early optimism has been tempered, however, by clinical results that have often been disappointing. It is clear that modifications of traditional methodologies to Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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promote endothelialization of vascular grafts are needed if endothelial cell transplantation is to be technically successful. What is even more important, however, is that academicians and physician researchers and investigators have grown to appreciate the complex biology of the cells forming the vascular wall, and that technically successful transplantation of endothelial cells does not necessarily presume a positive outcome clinically in terms of vascular graft performance. The simple hypotheses of early endothelial cell seeding studies have led to investigations providing enormous insights into the biology of vascular wall cells, as well as stimulating suggestions for alternative therapeutic avenues that may utilize concepts and techniques of endothelial cell seeding. The purpose of this chapter is to summarize the historic context and current base of knowledge regarding many of the technical issues relevant to endothelial cell seeding. Special reference will be made to recent exciting research methodologies designed to promote endothelial cell attachment to prosthetic graft materials—a technology known as electrostatic endothelial cell seeding. The wealth of animal studies has been reviewed elsewhere4 and will not be detailed here. Subsequent chapters in this text will summarize clinical outcomes. Historic Context The origin of the concept for the vascular prosthesis is credited to Voorhees, Jaretski, and Blakemore, who presented the first published hypothesis on synthetic tubes as replacements for natural blood vessel deficits in 1952.5 The first clinical applications of this hypothesis were published in 1954 and showed that the prosthetic tubes could be used in the arterial setting.6 This publication prompted a race to develop the best material for use as a prosthetic vascular graft. The two designs of large diameter vascular grafts (> 6 mm I.D.) which are used successfully today clinically are made of Dacron or expanded polytetrafluoroethylene (ePTFE). Patency rates for small diameter (< 6 mm I.D.) vascular prostheses configured from these same materials are unacceptable when utilized clinically, however, due to acute thrombus formation7,8 and chronic anastomotic hyperplasia9-11 which consequently prevent clinical usage. Malcolm Herring, M.D., introduced the hypothesis underlying the technology of endothelial cell seeding in 1978 in an attempt to increase the patency rate of small caliber prosthetic grafts. Herring proposed that if these small diameter vascular grafts could be utilized successfully, they could be of enormous clinical importance in peripheral and coronary bypass grafting procedures.12 Endothelial cell seeding since its inception has been largely experimental, with significant technical success in animal models. The technology has not been translated successfully to any extent thus far published in human vascular surgery. Surgeons have long expressed concerns about many of the technical issues which must be overcome in order to transplant endothelial cells successfully in a relevant time frame for human surgery. Moreover, the long term benefits of a transplanted endothelial cell lining on graft patency in patients has not been demonstrated. Many research efforts have focused upon methodologic issues in an effort to maximize cell harvest-

Tissue Engineering of Prosthetic Vascular Grafts

ing and transplantation efficiencies in an appropriate time frame for clinical practicality. We will review briefly the history of these studies and then describe the novel technology of electrostatic cell seeding, which we feel offers a solution to many of the deficiencies of other approaches for transplanting cells onto vascular prostheses. Autologous Endothelial Cell Harvesting Before endothelial cell seeding or transplantation can occur, autologous endothelial cells must be harvested from an “appropriate, autologous” source. Sources of endothelial cells most often reported have been: 1. Nonessential vessels (i.e., saphenous vein); and 2. Omental or subcutaneous adipose tissue (i.e., microvascular endothelial cells). From the nonessential vessels, endothelial cells have been harvested by two techniques: 1. Mechanical scraping;13 and 2. Enzymatic digestion.14,15 Mechanical scraping uses an abrasive action to remove the endothelial cells from the vascular wall, which leads to endothelial cell damage as well as the potential for significant contamination with smooth muscle cells.12,13 The overall efficiency of endothelial cell harvest from autologous vessels by mechanical methods is < 75%.14 Enzymatic harvesting of cells from the vascular tissue requires collagenase or trypsin digestion to remove the endothelium.16-18 This enzymatic digestion process can create damage to cellular proteins, which in turn affects cell viability and attachment potential to a vascular prosthetic surface.19,20 The overall efficiency of endothelial cell harvest from autologous vessels using the enzymatic harvesting technique varies between 80-100% of the available endothelial cells.14 Due to the limited availability of nonessential autologous vessels in patients, the ready supply of endothelial cells harvested by either of the previously mentioned techniques is necessarily limited. This limitation has prompted investigators to seek alternative sources which could supply a greater number of endothelial cells. The alternative sources which have been most intensively studied include the mesothelium and the microvasculature, each of which offers the potential for derivation of large cell numbers. Microvascular endothelial cells can be harvested from microvessels (arterioles, capillaries, and venules) found in adipose tissue.16,21,22 The overall efficiency of isolation of endothelial cells from microvessels utilizing enzymatic techniques has been reported to be approximately 84% ± 5%.23 Tissue culture of endothelial cells offers an alternative approach to derive large numbers of cells for transplantation from a small inoculum of harvested cells. The tissue culture approach was first reported by Graham et al.24 Jaffee et al17 initially reported that 92 h were required for an endothelial cell culture to double in number. However, improvements in cell culturing techniques have reduced that time to approximately 24 h using commercially available media (Clonetics Corporation. Unpublished data). Thus, relatively large numbers of endothelial cells can be obtained from low derived inocula of cells in a time span of days to weeks in culture. A number of valid issues with tissue culture have been voiced, however. The exposure of endothe-

Endothelial Cell Seeding: A Review

lial cells in culture to an undefined serum (e.g., fetal bovine serum) as a part of the tissue culturing media presents the opportunity for both genotypic and phenotypic modulation.25 In addition, multiple passages of cells in culture and media changes create vulnerability to cultures becoming contaminated. Endothelial Cell Seeding Techniques Numerous methodologies for endothelial cell transplantation onto prosthetic surfaces have been reported in the literature. These investigations can be differentiated and summarized by virtue of the unique physical force utilized in each transplantation process to seed endothelial cells to a vascular prosthetic blood-contacting surface: 1. Gravitational;12,26-29 2. Hydrostatic;30,31 and 3. More recently, electrostatic.32 The hypothesis for each of these approaches has been that the transplantation of endothelial cells (initially low surface density, cells/area) on a graft luminal surface will promote, through time, the development of a mature, physiologically appropriate, confluent graft luminal endothelial cell-lined neointima. The most basic of the three seeding techniques that have been tested utilizes gravitational forces to deliver endothelial cells to a vascular prosthetic luminal surface. The basic concept has been to fill the graft with the harvested endothelial cells resuspended in tissue culture medium or blood plasma. The filled graft is maintained horizontally and rotated periodically or continuously over a prolonged seeding time period. Biological glues have commonly been used with this technique to promote endothelial cell attachment to the prosthetic material. One disadvantage of using a glue is that any nonendothelialized surface becomes more thrombogenic due to the exposure of blood cells to the glue.8 A significant disadvantage of these early attempts at gravitational endothelial cell seeding was that the transplanted endothelial cells were in a spheroid morphology on the graft surface upon completion of the transplantation procedure. Thus, the potential existed for a significant loss of endothelial cells from the graft surface when blood flow was restored through the graft if the seeding time was short (< 2 h) and the cells had minimal time to flatten and mature. Endothelial cell losses up to 95% in the initial 24 h postimplantation were observed due to this cellular morphologic immaturity.33 To overcome this limitation of immature morphology, attempts were made to endothelialize the graft luminal surface by allowing adhesion to the graft with subsequent morphological maturation in tissue culture (7-14 days in vitro).27,28,34-37 These efforts did in fact result in minimal losses of endothelial cells from the grafts upon subsequent exposure to blood flow. However, as previously mentioned, the practicality of this approach remains of concern to surgeons and researchers. A long tissue culture period may not be acceptable for a patient who needs emergency bypass surgery. In addition, the questions of genotypic and/or phenotypic changes in cultured endothelial cells as well as the potential for infection (cell culture contamination) have been of major concern to investigators in this field.

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Hydrostatic seeding techniques use a pressure differential (internal pressure26,30 or external vacuum31) to force harvested endothelial cells resuspended in aliquots of heparinized autologous blood onto the luminal surface of microporous graft materials. Experience has revealed at least three major limitations of the hydrostatic seeding techniques. The first is that this technique is not adequate for fully heparinized patients due to the graft porosity. In heparinized patients, heparin interacts with the antithrombin III molecule which inhibits the action of thrombin to form subsequent clots within the graft pores, which is necessary to maintain hemostasis.38 Secondly, immediately following the preclotting technique, the surface may be rough and thrombogenic.26 The third limitation of this seeding technique is that, as previously mentioned, the endothelial cells adhere to the graft surface in a spheroid morphology (100% spheroid shaped) directly from the seeding suspension; thus an incubation period should be employed (> 2 h) for adhesion maturation prior to implantation.31 As described above, many investigators have attempted to increase the number of endothelial cells adhered to grafts as well as the magnitude of surface adhesion (flattening) by placing adhesive proteins such as fibronectin on the graft blood-contacting surface to act as a “glue”.39-48 Significant research efforts have focused on glue formulations including fibronectin, extracellular matrix, collagen, laminin, fibrin, fibroblast matrix, and plasma.27,40,41,43,45-50 The most commonly investigated “glue” has been fibronectin. Fibronectin is an adhesive glycoprotein which is found in the basement membrane to which the endothelial cells are attached in natural blood vessels. This glycoprotein is required to attach the endothelial cells to culture flasks in vitro and it has thus made sense to try it as the primary “glue“ to enhance endothelial cell adhesion to other artificial surfaces. The problem associated with using fibronectin or any other “glue” arises from the minimal number of endothelial cells which can be harvested relative to the total graft surface area to be covered, as well as the inefficiency of the seeding procedures.51 Any nonendothelialized graft surface or subsequent loss of endothelial cells from the surface upon implantation renders the exposed, fibronectin-treated graft surface attractive to platelets, thus potentially promoting thrombotic events that could lead to graft failure. 8 Ramalanjaona et al40,52 showed that the number of endothelial cells adhered was increased and the loss of cells upon exposure to shear stress was reduced using a fibronectin “glue”. The problem these investigators encountered was that the areas not endothelialized were thrombogenic, and experimental subjects required anticoagulant therapy to reduce the complications due to thrombus formation. One of the most difficult of the technical issues related to endothelial cell derivation from autologous sources and subsequent transplantation is the time required for these processes. It has been our experience that a minimum of 45-60 minutes is required for “immediate” seeding of endothelial cells onto a prosthetic graft being directly implanted in a patient. This minimum time frame does not allow for cell flattening and maturation onto the graft prior to restoration of blood flow through the graft. Some investigators have suggested that a time greater than two hours from

64

Tissue Engineering of Prosthetic Vascular Grafts

harvest to transplantation would most definitely not be acceptable clinically, due to the chance of genotypic and/or phenotypic changes to the endothelial cells which may occur while they are exposed to media containing undefined serum.25 In addition, it is obvious that longer time periods required for endothelial cell seeding of the vascular prostheses equates to longer periods of time for patients under anesthesia, and thus the potential for complications. The length of time and dosage of anesthesia and its safety are difficult to generalize due to the fact that exposure complication encompasses multiple drugs, techniques, anesthetists, and patients.53,54 The necessity for an incubation period for attachment and morphological maturation of transplanted endothelial cells is related to the nature of the electrostatic interactions between the endothelial cells and the prosthetic graft materials. Endothelial cells are negatively charged.55,56 The clinically successful vascular prosthetics (i.e., ePTFE) are highly negatively charged. This negativity repels platelets which are also negatively charged.57-61 Thus, the initial adherence of transplanted endothelial cells must overcome this negativenegative charge repulsive force (long range) between cells and graft material in order for the seeding procedure to be successful. Experiments using platelets have demonstrated that the cellular adhesion of platelets on a negatively charged substrate is one order of magnitude less (ten times) than expected by gravitational settling alone due to this electrostatic, repulsive, interaction which must be overcome by a stochastic process.62-64 Similar short range repulsive interactions alter cellular morphologies by preventing or slowing morphological maturation even when cells overcome the repulsive interactions with the material and attach to the surface. This reality is further demonstrated by endothelial cell and fibroblast studies that have shown the dependence of cell adherence on the substrate surface charge.65-70 These studies used varying substrates with varying surface charges to study cell adhesion, spreading, and contact regions between the cells and the substrates. The overall results from

these studies indicated that an increasingly positively charged surface leads to enhanced adhesion, spreading, and magnitude of contact regions. The results on increasingly negatively charged substrates indicated the inverse, with inhibited adhesion, reduced spreading, and reduced contact regions. Electrostatic Endothelial Cell Seeding The only conclusion that can be drawn from evaluating 15 years of research efforts related to cell seeding techniques is that few concrete technical advancements have been made. We have recently proposed and tested a novel device that we feel will be of significant value in improving the efficiency of transplanted cell attachment, as well as minimizing cellular losses upon implantation.32 The technique is called electrostatic endothelial cell seeding. The electrostatic seeding technique has been evaluated in vitro using the prototype apparatus shown in Figure 3.1. The key to this technique is that it enhances endothelial cell adhesion by inducing a temporary positive surface charge or a “temporary glue” on the negatively charged ePTFE graft luminal surface. Following cell transplantation the ePTFE graft luminal surface reverts to its original highly negatively charged surface. Thus, any nonendothelialized graft surfaces or any exposed graft surfaces resultant from endothelial cell losses upon restoration of blood flow remain nonthrombogenic due to the restored high negative surface charge of the graft material itself. The basic issue underlying electrostatic endothelial cell seeding is: “How can the surface potential of the graft be altered to attract endothelial cells without rendering the surface thrombogenic?” The electrostatic endothelial cell seeding technique takes advantage of graft material properties (dielectric material). When a dielectric material is placed within a capacitor (electrostatic seeding apparatus), the electrons of the atoms and ions which make up the dielectric material (near surface) are attracted to the capacitor surface which has accumulated the positive charge. The nuclei of the dielectric material (near surface) are attracted to the

Fig. 3.1. Prototype electrostatic endothelial cell seeding apparatus. Features include: (a) external conductor, (b) internal conductor, (c) electric motor drive/pulley system, (d) filling apparatus, (e) voltage source, (f ) pillow blocks, and (g) internal conductor end supports.

d b

c

g

e a f

Endothelial Cell Seeding: A Review

negatively charged surface. These small displacements, or polarizations, are what induce the surface charge on the graft luminal surface. It should be noted that the electrons in a dielectric material are not free and the displacements of the electrons are very slight. Also, the interior volume of the graft material, dielectric, remains unchanged, thus leaving a net charge of zero over the dielectric material.71,72 Several in vitro studies have been performed utilizing this electrostatic seeding technique. When human umbilical vein endothelial cells were transplanted onto 4 mm I.D. ePTFE using the electrostatic cell seeding technique, a complete nodal area coverage of morphologically mature (completely flattened) endothelial cells (73,540 cells/cm2) was obtained in 16 minutes (+1.0 Volt applied to apparatus) with minimal cellular membrane damage or effect on endothelial cell viability.32 A section of electrostatically seeded ePTFE is illustrated in the scanning electron micrograph in Figure 3.2. These in vitro evaluations of the electrostatic seeding technique revealed no significant losses of endothelial cells upon exposure of the graft to a wall shear stress of 15 dynes/cm2 for up to 120 minutes immediately after seeding. The majority of endothelial cell loss (up to 30%) occurs within the first 30 minutes of implantation using traditional techniques.33,40 Thus, the electrostatic seeding procedure is superior to the gravitational and hydrostatic seeding procedures in terms of the seeding time required, magnitude of endothelial cell adhesion (attachment), and cellular retention. It is also speculated that the actin microfilaments which make up the endothelial cell cytoskeleton and possess an electrostatic potential will also be rearranged by the electrostatic endothelial cell seeding technique to assist in the maintenance of endothelial cell adhesion (act as an anchoring system), although this has not yet been demonstrated in our experiments.73,74 It is also hoped that the enhanced cell attachment resultant from electrostatic endothelial cell seeding may allow the seeded endothelial cells to synthesize the

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necessary fibronectin (< 3 days) and basement membrane collagen (< 1 week) to maintain cellular adhesion on the graft for the long term.75-77 Preliminary in vivo studies (unpublished) using a canine femoral artery implantation model suggests thromboresistance of the electrostatically seeded grafts. Genetically Engineering Endothelial Cells for Seeding Vascular Prostheses It has recently been suggested that genetically engineered endothelial cells could be transplanted, which might promote the reduction/prevention of anastomotic hyperplasia and enhance thromboresistance of small diameter vascular prostheses. Additional theoretical therapeutic applications using manipulated endothelial cells include the possibilities of gene replacement, correction, or augmentation which may be effective in the treatment of atherosclerosis. Localized drug delivery may also be feasible due to the direct contact of transplanted endothelial cells within the blood stream.78,79 The initial use of transplanted, genetically modified endothelial cells is credited to Zweibel et al.80,81 Since that time, the study of endothelial cell transfection (incorporation of foreign DNA) using retroviral and electroporation methodologies has demonstrated long term cell viability as well as stable transfection using the reporter genes β-galactosidase, chloramphenicol acetyltransferase, luciferase and human tissue plasminogen activator (tPA), growth hormone, urokinase, and nitric oxide synthase genes.82-93 The efficiencies of retroviral and electroporation transfection of endothelial cells are 1-2% and 10%, respectively.85,89 Adenovirusmediated gene transfer is highly dose dependent and produces only transient transfection. 87 Transfection and endothelial cell seeding are procedures each of which requires long time periods of laboratory work, which again raises a red flag among the pragmatists concerned with

Fig. 3.2. Scanning electron micrograph immediately after the electrostatic seeding of human umbilical vein endothelial cells (+1.0 Volt applied for 16 minutes) on ePTFE (GORE-TEX®; 30 µm internodal distance) illustrating the complete, morphologically mature coverage of the nodal areas (Magnification x750).

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additional genotypic (other than desired alteration) or phenotypic changes and infections. Transplantation of transfected endothelial cells does not appear to affect initial seeding densities.82 However, retention of endothelial cells is significantly reduced due to retroviral transfection (no data available for electroporated cells). This decrease in retention is in addition to the already expected loss using a gravitational seeding procedure. Thus, further refinement of all of these techniques is imperative in order for them to be embraced clinically. It is speculated that the electrostatic endothelial cell seeding device previously described may be used to successfully seed/transfect endothelial cells simultaneously in 16 minutes. The concept is being evaluated by applying high voltage pulses (> 200 Volts) while performing the low voltage, +1.0 Volt, seeding procedure. The high voltage pulses (100 µs duration) are either a single pulse or multiple (i.e., 4) pulses at set time intervals (4-5 seconds). It is expected that the electroporation pulses will be applied midway through the seeding procedure when the endothelial cells and DNA will be in close proximity at the graft luminal surface. It is expected that the small electric field used for seeding (and maintained continuously) will allow increased cellular transfection by maintaining the cellular membrane pores created for a slightly longer time period. This speculation is based on the theory that the cell membrane pores are created by protrusions of cellular membrane structures (and/or underlying cortical layer) induced by application of an electric field.94 This, in conjunction with the electrophoretic motion of the DNA towards the graft luminal surface and the endothelial cells, is expected to contribute to an increased transfection efficiency. Currently, electroporation techniques have a maximum overall transfection efficiency of 10-20%.95,96 Along with the increase in transfection efficiency, the electrostatic seeding/transfection procedure may also contribute to increased or maintained cellular retentions upon implantation when compared to currently available techniques. Conclusion A wealth of information has been derived from research protocols investigating the potential of endothelial cell seeding in improving small diameter vascular graft function. Realization of the hypothesized potential of endothelial cell seeding for patients has remained heretofore elusive. This may be due in large measure to the complex biology of the vascular wall and the disruption created by the transplantation process. One wonders, however, whether or not that potential has yet been adequately tested, as various traditional methodologies and techniques for endothelial cell transplantation have all been, to one extent or another, inadequate when applied clinically. The technique of electrostatic endothelial cell seeding which we have described in this chapter in the historic context of endothelial cell seeding offers the potential to circumvent or alleviate many of those technical issues. It is with renewed enthusiasm then that we are revisiting the technology of endothelial cell seeding of vascular grafts armed with new technical information and new insights into the biology of vascular cells.

Tissue Engineering of Prosthetic Vascular Grafts

References 1. Berger K, Sauvage LR, Rao AA, Wood S. Healing of arterial prostheses in man: Its incompleteness. Ann Surg 1972; 175:119-27. 2. Sauvage LR, Berger KE, Wood SJ et al. Interspecies heal of porous arterial prostheses. Arch Surg 1974; 109:698-706. 3. Clowes AW, Kirkman TR, Reidy MA. Mechanisms of arterial graft healing. Rapid transmural capillary ingrowth provides a source of intimal endothelium and smooth muscle in porous PTFE prostheses. Am J Pathol 1986; 123:221-9. 4. Schmidt SP, Meerbaum SO, Sharp WV. Endothelial-lined vascular prostheses. In: Wise DL, Trantolo DJ, Altobelli DE et al., eds. Encyclopedic handbook of biomaterials and bioengineering, Part B: Applications. New York: Marcel Dekker, Inc., 1995:1089-1110. 5. Voorhees AB, Jaretski A, Blakemore AH. Use of tubes constructed from vinyon ‘n’ cloth in bridging arterial deficits. Ann Surg 1952; 135:332-8. 6. Blakemore A, Voorhees AB. The use of tubes constructed of vinyon ‘n’ cloth in bridging arterial deficits: Experimental and clinical. Ann Surg 1954; 140:324-30. 7. Clagett GP, Burkel WE, Sharefkin JB et al. Platelet activity in vivo in dogs with arterial prostheses seeded with endothelial cells. Circulation 1984; 69:632-9. 8. Kempczinski RF, Ramalanjaona GR, Douville C et al. Thrombogenicity of a fibronectin-coated, experimental polytetrafluoroethylene graft. Surgery 1987; 101:439-44. 9. Clowes AW, Reidy MA. Mechanisms of arterial graft failure: The role of cellular proliferation. Ann NY Acad Sci 1987; 516:673-8. 10. Clowes AW, Kirkman TR, Clowes MM. Mechanisms of arterial graft failure. II. Chronic endothelial cell and smooth muscle cell proliferation in healing polytetrafluoroethylene prostheses. J Vasc Surg 1986; 3:877-84. 11. Clowes AW, Gown AM, Hanson SR et al. Mechanism of arterial graft failure I: Role of cellular proliferation in early healing of PTFE grafts. Am J Path 1985; 118:43-54. 12. Herring M, Gardner A, Glover J. A single-staged technique for seeding vascular grafts with autogenous endothelium. Surgery 1978; 84:848-54. 13. Herring M, Gardner A, Glover J. Seeding human arterial prostheses with mechanically derived endothelium. J Vasc Surg 1984; 1:279-89. 14. Stanley JC, Graham LM, Burkel WE. Endothelial cell seeded synthetic vascular grafts. In: Vascular Graft Update: Safety and Performance. Philadelphia: ASTM, 1986:33-43. 15. Graham LM, Burkel WE, Ford JW et al. Expanded polytetrafluoroethylene vascular prostheses seeded with enzymatically derived and cultured canine endothelial cells. Surgery 1982; 91:550-9. 16. Sharp WV, Schmidt SP, Meerbaum SO et al. Derivation of human microvascular endothelial cells for prosthetic vascular graft seeding. Ann Vasc Surg 1989; 3:104-7. 17. Jaffe EA, Nachman R, Becker C et al. Culture of human endothelial cells derived from umbilical veins. Identification by morphologic and immunologic criteria. J Clin Invest 1973; 52:2745-56. 18. Bourke BM, Roche WR, Appleberg M. Endothelial cell harvest for seeding vascular prostheses: The influence of technique on cell function, viability, and number. J Vasc Surg 1986; 4:257-63.

Endothelial Cell Seeding: A Review 19. Gimbrone MA, Cotran RS, Folkman J. Human vascular endothelial cells in culture. Growth and DNA synthesis. J Cell Biol 1974; 60:673-84. 20. Sharefkin JB, Van Wart HE, Cruess DF et al. Adult human endothelial cell enzymatic harvesting. Estimates of efficiency and comparisons of crude and partially purified bacterial collagenase preparations by replicate microwell culture and fibronectin degradation measured by enzyme-linked immunosorbent assay. J Vasc Surg 1986; 4:567-77. 21. Jarrell BE, Williams SK, Stokes G et al. Use of freshly isolated capillary endothelial cells for the immediate establishment of a monolayer on a vascular graft at surgery. Surgery 1986; 100:392-9. 22. Williams SK, Wang TF, Castrillo R et al. Liposuction-derived human fat used for vascular graft sodding contains endothelial cells and not mesothelial cells as the major cell type. J Vasc Surg 1994; 19:916-23. 23. Rupnick MA, Hubbard FA, Pratt K et al. Endothelialization of vascular prosthetic surfaces after seeding or sodding with human microvascular endothelial cells. J Vasc Surg 1989; 9:788-95. 24. Graham LM, Vinter DW, Ford JW et al. Endothelial cell seeding of prosthetic vascular grafts. Early experimental studies with cultured autologous canine endothelium. Arch Surg 1980; 115:929-33. 25. Schmidt SP, Sharp WV, Evancho MM et al. Endothelial cell seeding of prosthetic vascular grafts-current status. In: Szycher M, ed. High Performance Biomaterials. Lancaster: Technomic Publishing Co., 1991:483-96. 26. Yates SG, Barros AAB, Berger K et al. The preclotting of porous arterial prostheses. Ann Surg 1978; 188:611-22. 27. Anderson JS, Price TM, Hanson SR et al. In vitro endothelialization of small-caliber vascular grafts. Surgery 1987; 101:577-86. 28. Foxall TL, Auger KR, Callow AD et al. Adult human endothelial cell coverage of small-caliber dacron and polytetrafluoroethylene vascular prostheses in vitro. J Surg Res 1986; 41:158-72. 29. Mazzucotelli JP, Roudiere JL, Bernex F et al. A new device for endothelial cell seeding of a small-caliber vascular prosthesis. Artif Organs 1993; 17:787-90. 30. Kempczinski RF, Rosenman JE, Pearce WH et al. Endothelial cell seeding of a new PTFE vascular prosthesis. J Vasc Surg 1985; 2:424-9. 31. Van Wachem PB, Stronck JWS, Koers-Zuideveld R et al. Vacuum cell seeding: A new method for the fast application of an evenly distributed cell layer on porous vascular grafts. Biomaterials 1990; 11:602-6. 32. Bowlin GL. Electrostatic endothelial cell seeding of vascular prostheses. Doctoral Dissertation, The University of Akron, 1996:1-354. 33. Rosenman JE, Kempczinski RF, Pearce WH et al. Kinetics of endothelial cell seeding. J Vasc Surg 1985; 2:778-84. 34. Shindo S, Takagi A, Whittemore AD. Improved patency of collagen-impregnated grafts after in vitro autogenous endothelial cell seeding. J Vasc Surg 1987; 6:325-32. 35. Prendiville EJ, Coleman JE, Callow AD et al. Increased in vitro incubation time of endothelial cells on fibronectintreated ePTFE increases cell retention in blood flow. Eur J Vasc Surg 1991; 5:311-9. 36. Sentissi JM, Ramberg K, O’Donnell K et al. The effect of flow on vascular endothelial cells grown in tissue culture on polytetrafluoroethylene grafts. Surgery 1986; 99:337-43.

67 37. Budd JS, Allen KE, Hartley G et al. The effect of preformed confluent endothelial cell monolayers on the patency and thrombogenicity of small calibre vascular grafts. Eur J Vasc Surg 1991; 5:397-405. 38. Turgeon ML. Clinical Hematology: Theory and Procedure. Boston: Little, Brown, and Co., 1988:279-86. 39. Kesler KA, Herring MB, Arnold MP et al. Enhanced strength of endothelial attachment on polyester elastomer and polytetrafluoroethylene graft surfaces with fibronectin substrate. J Vasc Surg 1986; 3:58-64. 40. Ramalanjaona G, Kempczinski RF, Rosenman JE et al. The effect of fibronectin coating on endothelial cell kinetics on polytetrafluoroethylene grafts. J Vasc Surg 1986; 3:264-72. 41. Seeger JM, Klingman N. Improved in vivo endothelialization of prosthetic grafts by surface modification with fibronectin. J Vasc Surg 1988; 8:476-82. 42. Pratt KJ, Jarrell BE, Williams SK et al. Kinetics of endothelial cell-surface attachment forces. J Vasc Surg 1988; 7:591-9. 43. Van Wachem PB, Vreriks CM, Beugeling T et al. The influence of protein adsorption on interactions of cultured human endothelial cells with polymers. J Biomed Mater Res 1987; 21:701-18. 44. Van Wachem PB, Beuqeling T, Mallens BW et al. Deposition of endothelial fibronectin on polymeric surfaces. Biomaterials 1988; 9:121-3. 45. Schneider A, Melmed RN, Schwalb H et al. An improved method for endothelial cell seeding on polytetrafluoroethylene small caliber vascular grafts. J Vasc Surg 1992; 15:649-56. 46. Lee YS, Park DK, Kim YB et al. Endothelial cell seeding onto the extracellular matrix of fibroblasts for the development of a small diameter polyurethane vessel. ASAIO J 1993; 39: M740-5. 47. Bellon JM, Bujan J, Honduvilla NG et al. Endothelial cell seeding of polytetrafluoroethylene vascular prostheses coated with a fibroblastic matrix. Ann Vasc Surg 1993; 7:549-55. 48. Gosselin C, Vorp DA, Warty V et al. ePTFE coating with fibrin glue, FGF-1, and heparin: Effect on retention of seeded endothelial cells. J Surg Res 1996; 60:327-332. 49. Williams SK, Jarrell BE, Friend L et al. Adult human endothelial cell compatibility with prosthetic graft material. J Surg Res 1985; 38:618-29. 50. Kaehler J, Zilla P, Fasol R et al. Precoating substrate and surface configuration determine adherence and spreading of seeded endothelial cells on polytetrafluoroethylene grafts. J Vasc Surg 1989; 9:535-41. 51. Lindblad B, Burkel WE, Wakefield TW et al. Endothelial cell seeding efficiency onto expanded polytetrafluoroethylene grafts with different coatings. ACTA Chir Scand 1986; 152:653-6. 52. Ramalanjaona GR, Kempczinski RF, Ogle JD et al. Fibronectin coating of an experimental PTFE vascular prosthesis. J Surg Res 1986; 41:479-83. 53. Longnecker DE, Murphy FL. Introduction to anesthesia. 8th ed., Philadelphia: W.B. Saunders Co., Harcourt Brace Jovanovich, Inc., 1992:419-27. 54. Guyton AC. Textbook of medical physiology. 8th ed. Philadelphia: W.B. Saunders Co., Harcourt Brace Jovanovich, Inc., 1991:269-71. 55. Sawyer PN, Harshaw DH. Electroosmotic characteristics of canine aorta and vena cava walls. Biophy J 1966; 6:653-63.

68 56. Srinivasan S, Sawyer PN. Role of surface charge of the blood vessel wall, blood cells, and prosthetic material in intravascular thrombosis. J Colloid Sci 1970; 32:456-63. 57. Sawyer PN, Srinivasan S. Studies on the biophysics of intravascular thrombosis. Am J Surg 967; 114:42-59. 58. Lowell J. The electrification of polymers by metals. J Phys D: Appl Phys 1976; 9:1571-85. 59. Yu ZZ, Watson PK, Facci JS. The contact charging of PTFE by mercury: The effect of a thiophene monolayer on charge exchange. J Phys D: Appl Phys 1990; 23:1207-11. 60. Yu ZZ, Keith PK. Contact charge accumulation and reversal on polystyrene and PTFE films upon repeated contacts with mercury. J Phys D: Appl Phys 1989; 22:798-801. 61. Abramson HA. The electrophoresis of the blood platelets of the horse with the reference to their origin and to thrombus formation. J Exp Med 1928; 47:677-83. 62. Marmur A, Gill WN, Ruckenstein E. Kinetics of cell deposition under action of an external force. Bull Math Biol 1976; 38:713-21. 63. Ruckenstein E, Prieve DC. Dynamics of cell deposition on surfaces. J Theor Biol 1975; 51:429-38. 64. Ruckenstein E, Marmur A, Rakower SR. Sedimentation and adhesion of platelets onto horizontal glass surface. Thrombus Haemostas 1976; 36:334-42. 65. Van Wachem PB, Hogt AH, Beugeling T et al. Adhesion of cultured human endothelial cells onto methacrylate polymers with varying surface wettability and charge. Biomaterials 1987; 8:322-9. 66. Macarak EJ, Howard PS. Adhesion of endothelial cells to extracellular matrix proteins. J Cell Phys 1983; 116:76-86. 67. Sugimoto Y. Effects on the adhesion and locomotion of mouse fibroblasts by their interacting with differently charged substrates. Exp Cell Res 1981; 135:39-45. 68. Schakenraad JM, Arends J, Busscher HJ et al. Kinetics of cell spreading on protein precoated substrata: A study of interfacial aspects. Biomaterials 1989; 10:43-50. 69. Niu S, Matsuda T, Oka T. Endothelialization on various segmented polyurethanes: Cellular behavior and its substrate dependency. ASAIO Trans 1990; 36: M164-8. 70. Van Wachem PB, Schakenraad JM, Feijen J et al. Adhesion and spreading of cultured endothelial cells on modified and unmodified poly(ethyleneterephthalate): A morphological study. Biomaterials 1989; 10:532-9. 71. Fink DG, Christiansen D, eds. Electronics engineers’ handbook. 3rd ed. New York: McGraw-Hill Book Co., 1989:6:29-64. 72. Halliday D, Resnick R. Fundamentals of physics. 2nd ed. New York: John Wiley & Sons, Inc., 1981:484-502. 73. Ando T, Kobayashi N, Munekata E. Electrostatic potential around actin. In: Sugi H, Pollack GH, eds. Mechanism of myofilament sliding in muscle contraction. New York: Plenum Press, 1993:361-76. 74. Gottlieb AI, Langille BL, Wong MKK et al. Biology of disease: Structure and function of the endothelial cytoskelaton. Lab Invest 1991; 65:123-37. 75. Jaffe EA, Mosher DF. Synthesis of fibronectin by cultured human endothelial cells. J Exp Med 1978; 147:1779-91. 76. Jaffe EA, Minick CR, Adelman B et al. Synthesis of basement membrane collagen by cultured human endothelial cells. J Exp Med 1976; 144:209-25. 77. Howard BV, Macarak EJ, Gunson D et al. Characterization of the collagen synthesized by endothelial cells in culture. Proc Natl Acad Sci USA 1976; 73:2361-4.

Tissue Engineering of Prosthetic Vascular Grafts 78. Nabel EG, Plautz G, Nabel GJ. Gene transfer into vascular cells. JACC 1991; 17:189B-94B. 79. Callow AD. The vascular endothelial cell as a vehicle for gene therapy. J Vasc Surg 1990; 11:793-8. 80. Zweibel JA, Freeman SM, Kantoff PW et al. High-level recombinant gene expression in rabbit endothelial cells transduced by retroviral vectors. Science 1989; 243:220-2. 81. Ryan US, Hayes BA, Maxwell G et al. Endothelial cells as vehicles for continuous release of recombinant gene products. In: Zilla P, Fasol R, Callow A, eds. Applied cardiovascular biology. Basel: S. Karger, 1990:22-9. 82. Sackman JE, Freeman MB, Petersen MG et al. Synthetic vascular grafts seeded with genetically modified endothelium in the dog: Evaluation of the effect of seeding technique and retroviral vector on cell persistence in vivo. Cell Trans 1995; 4:219-35. 83. Huber TS, Welling TH, Sarkar R et al. Effects of retroviralmediated tissue plasminogen activator gene transfer and expression on adherence and proliferation of canine endothelial cells seeded onto expanded polytetrafluoroethylene. J Vasc Surg 1995; 22:795-803. 84. Powell JT, Klaasse Bos JM, Van Mourik JA. The uptake and expression of the factor VIII and reporter genes by vascular cells. FEBES 1992; 303:173-7. 85. Kahn ML, Lee SW, Dichek DA. Optimization of retroviral vector-mediated gene transfer into endothelial cells in vitro. Circ Res 1992; 71:1508-17. 86. Hong Z, Guangbin Z, Xiaojun Z et al. Enhanced adenoassociated virus vector expression by adenovirus proteincationic liposome complex. Chinese Med J 1995; 108:332-7. 87. Hong Z, Guangbin Z, Airu Z et al. Adenovirus mediated gene transfer of vascular smooth muscle cells and endothelial cells in vitro. Chinese Med J 1995; 108:493-6. 88. Von der Leyen HE, Gibbons GH, Morishita R et al. Gene therapy inhibiting neointimal vascular lesions: In vivo transfer of endothelial cell nitric oxide synthase gene. Proc Natl Acad Sci 1995; 92:1137-41. 89. Kotnis RA, Thompson MM, Eady SL et al. Attachment, replication and thrombogenecity of genetically modified endothelial cells. Eur J Vasc Endovasc Surg 1995; 9:335-40. 90. Wilson JM, Birinyi LK, Salomon RN et al. Implantation of vascular grafts lined with genetically modified endothelial cells. Science 1989; 244:1344-6. 91. Dunn PF, Newman KD, Jones M et al. Seeding of vascular grafts with genetically modified endothelial cells: Secretion of recombinant TPA results in decreased seeded cell retention in vitro and in vivo. Circulation 1996; 93:1439-46. 92. Schwachtgen JL, Ferreira V, Meyer D et al. Optimization of the transfection of human endothelial cells by electroporation. Biotechiques 1994; 5:882-7. 93. Kotnis RA, Thompson MM, Eady SL et al. Optimisation of gene transfer into vascular endothelial cells using electroporation. Eur J Vasc Endovasc Surg 1995; 9:71-9 94. Popov SV, Margolis LB. Formation of cell outgrowths by external force: A model study. J Cell Sci 1988; 90:379-89. 95. Spencer SC. Electroporation technique of DNA transfection. In: Murray EJ, ed. Methods in molecular biology. Clifton: The Humana Press, Inc., 1991:45-52. 96. Davis LG, Dibner MD, Battey JF. Basic methods in molecular biology. Norwalk: Appleton & Lange, 1986:293-95.

CHAPTER 4 Surface Precoating in the 1980s: A First Taste of Cell-Matrix Interactions J. Vincent Smyth, Michael G. Walker

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echnical difficulties and the frequent lack of available saphenous vein for peripheral arterial reconstruction resulted in the development of a variety of prosthetic materials from which grafts could be constructed. Of these, polyester elastomer (Dacron) and expanded polytetrafluorethylene (ePTFE) emerged as the most successful in clinical practice, and remain leaders in the field today. However, small caliber prosthetic grafts have consistently shown lower patency rates, especially when passing across the knee joint, when compared with autologous vein conduits. The understanding that the endothelial lining of vein grafts was responsible for the improved patency rates achieved in peripheral bypass naturally led to attempts to create such a lining on prosthetic grafts, when explants showed that no significant endothelialization of such grafts occurs naturally in humans, contrary to the results seen in animals where virtually complete endothelial linings develop on prosthetic grafts in a matter of weeks. Almost all of the work performed up to 1980 utilized animal tissue models such as bovine, canine or murine endothelium, and extremely promising results had been repeatedly achieved both in vivo and in vitro in terms of reduction in graft thrombogenicity, resistance to infection, accelerated re-endothelialization and improved patency.1-6 In order to make the transition to clinical practice, however, it was clear that human cell lines would be required, and human umbilical vein endothelial cells (Huvecs) were used as the principal experimental model until Jarrell showed that human adult endothelial cells (HAECs) could be harvested from saphenous veins and cultured.7 This work has now advanced so that the acute enzymatic harvesting of autologous vein can be routinely achieved, with virtually complete denudation of the donor vein segment and a large proportion of harvested cells remaining viable for culture. There is now extensive experience with both cell lines toward the desired end point of successful graft seeding to the rapid development of an endothelial monolayer that resists the shear stress of flow.8 Almost as soon as the successful culture of Huvecs on prosthetic graft material was first reported, it was clear that human cell lines were more fragile and fastidious in their requirements than the canine or bovine endothelial cultures that had principally been used up to this point. The establishment of a culture was slower even under optimal conditions, and when achieved these were less hardy and prone to failure from infection. Even so, successful culture was reported on all the common graft materials, with typical appearances of cobblestone morphology and cells expressing endothelium-specific markers such as vWF. Although human cell lines showed slower cell division and spreading across the culture surfaces than animal cell lines, even under the optimal experimental conditions of tissue Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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culture, the main reason for the delayed development of the monolayer was found to be low cell attachment to the graft.

Endothelial Cell Attachment to Prosthetic Grafts The first seeding experiments were carried out in what we would now classify as a static system, where the cells were introduced to the graft surface in a suspension of culture medium and allowed to settle on the graft over a period of time, then cultured in situ. Early work relied on qualitative phase contrast microscopy of the seeded surface at intervals after seeding to assess degree of coverage. Later, the work of Sharefkin led to the use of the radio-isotope 111Indium as an endothelial cell label,9 and attachment to grafts could be accurately quantified. It was discovered that prosthetic grafts exhibit a uniformly poor surface for retention of seeded cells, typically around 4% for ePTFE,10,11 while cell adherence to plain Dacron was negligible because of its porosity, resulting in leakage of the seeding solution.12,13 Three approaches to the problem of rapidly developing a monolayer were explored. The proportion of seeded cells that remained attached to the graft could be increased by increasing the length of time that the seeding suspension was exposed to the graft material.10,14-16 However, the seeding kinetics indicated that the rate of increase fell off after around 30 minutes, and that even with exposure times of over 60 minutes, the development of a confluent monolayer was not significantly accelerated. Additionally, prolonged incubation times are incompatible with the constraints of the operating room. Using more endothelial cells per graft area in the seeding solution, a higher seeding density, produced a greater number of cells attaching to the graft, but only in the same proportion and consequently with an enormous waste of seeded cells.15,17 At established attachment rates to plain

Fig. 4.1. Cell attachment to prosthetic grafts as a function of graft coating.

Tissue Engineering of Prosthetic Vascular Grafts

grafts, up to 20 times as many cells as were required to form a monolayer would be needed for seeding, thus greatly increasing the required cell numbers. Although this could be accommodated by the use of preculture in the laboratory,18 this method is not compatible with routine surgical practice and would necessitate a preliminary operation for vein harvest and carry risks of disease transmission or graft infection. To overcome the problem of limited cell harvests, the use of microvascular (capillary) endothelium19 and mesothelial20 cells have been proposed, but these only offer a partial solution in terms of available cell numbers, as they share many of the attachment and adhesion difficulties experienced with Huvecs and HAECs, and raise additional problems with tissue availability. As a consequence, to achieve rapid development of confluence in the clinical arena where such large cell numbers could not be obtained acutely using established methods, either new sources of endothelial cells would be required or cell attachment significantly improved to permit successful seeding at subconfluent densities. It had been noticed that the results of human cell seeding work did not parallel those seen in the animal in vivo models of prosthetic graft seeding. Significantly better attachment was seen in the latter, and it was realized that the preclot typically used for the endothelial cell suspension acted as a graft coating, improving cell adherence. By using preclot or serum to coat the graft prior to seeding the endothelial cells, attachment rates could be dramatically improved to over 60% of seeded cells in a range of models.10,21 As a result, investigators began precoating the culture dishes and seeding surfaces with a variety of matrices before the cells were exposed, and similar improvements in cell attachment were seen11,22-27 (Figs. 4.1-4.3). Comparative studies of a range of potential graft surface coatings were made. Because of Dacron’s inherent porosity, gelatin or collagensealed grafts were frequently used, and this not surprisingly

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Fig. 4.2. Electron micrograph of uncoated prosthetic graft.

Fig. 4.3. Electron micrograph of fibronectin-coated graft after endothelial seeding.

improved initial cell adherence accordingly, as did all graft coating matrices compared to the plain graft materials. Histological examination of the vascular endothelium indicated that in the body it was supported by a proteinrich basement membrane layer, and that this layer was produced actively at least in part by the endothelium. Examination of endothelial cells in culture showed that the adherent cells secreted a similar layer on artificial surfaces28,29 and that a principal component of the basement membrane was fibronectin (FN). Following the work of Ramalanjoana21 and others,30-32 FN has become the most commonly used seeding matrix because of efficacy and availability, giving consistently high cell adherences in experiments using a variety of graft types, including ePTFE and Dacron. Preclot, serum, plasma, laminin, gelatin, collagen and fibrin glue have also been investigated,10,27,33-36 the first three because they are readily available in the clinical situation, and the remainder because they, along with fibronectin, are found as native constituents of extracellular matrix. Although preclot, serum and

plasma show good rates of endothelial attachment, this is thought to be due to their high FN content. Laminin and collagens I, III and IV have not proved as effective as FN in facilitating endothelial cell attachment. Gelatin-sealed Dacron grafts show good initial cell adherence, but the stability of the gelatin crosslinking under physiological conditions is uncertain, as it degrades and may not provide a suitable matrix for long enough to allow the cells to establish their own FN-rich basement membrane layer on the graft. Fibrin glue has been used successfully in some seeding work, but requires fibrinolytic inhibition prior to use without offering any advantages in terms of cell retention. In the laboratory model, however, almost equivalent results to those with FN have been achieved and, in retrospect, it is clear that one reason for the success of the early animal seeding experiments using prosthetic grafts had been the serendipitous use of FN-rich preclot as a readily available cell suspension to seed the graft, sealing the porous Dacron used in this work.

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Many workers have investigated the attachment kinetics of endothelial cells to FN-coated graft surfaces.10,15,27,37 In the laboratory, the optimum incubation time for cell attachment to a coating matrix appears to be 30 minutes with no significant benefit evident beyond this, though some workers have found as little as 15 minutes adequate. As with uncoated grafts, improved cell attachment is also seen with increasing the seeding density, up to a point whereafter the proportion of seeded cells attaching falls off, though the absolute number continues to increase. The kinetics of improved cell adherence to FN-coated graft surfaces implied some specific binding mechanism, and this was confirmed by the finding that endothelial cells express specific FN receptors on their cell membrane. Comparative studies of different grafts seem to suggest that ePTFE is a better surface for matrix coating than Dacron, although there is some disagreement between individual studies. This may be related to the nature of the graft surface at the fiber level, with a flatter surface being offered by the multinodal fibrillary construction than the woven or knitted Dacron. Minor refinements in terms of graft composition, pore size, internodal distances or woven vs. knitted Dacron, do not seem to affect the cell seeding kinetics significantly. On precoated grafts, the differences in endothelial cell adherence between graft materials is not striking once the porosity of the Dacron is reduced by sealing the graft with the surface coating, and this can be explained by the production of essentially similar attachment characteristics at the molecular level on the surfaces to which the seeded endothelial cells are exposed. With the establishment of a FN-rich coating as the optimal seeding surface, several workers have investigated the adsorption kinetics of FN onto prosthetic grafts,38,39 showing that the development of FN-graft binding is proportional to the FN concentration and the duration of incubation, though there may be minor differences between graft materials. The optimum concentration of FN to use as the seeding matrix has been investigated,40 and the relationship between the concentration of FN graft coating and rates of endothelial cell adherence approximates first order kinetics to 50 µg/ml, with no benefit obtained by increasing this to 250 µg/ml, further evidence for specific FN-binding sites on the cell surface.

Cell Migration It had been noted that following seeding, cells initially attached to the graft in a rounded morphology, later developing a more flattened phenotype responsible for the ‘cobblestone’ appearance of the mature endothelial monolayer.32 Increased duration of incubation was associated with progression of the cell morphology from rounded and poorly attached, to flattened and adherent, and this was generally accepted to indicate the acclimatization of the seeded cell to the seeding surface. The speedier development of an adherent monolayer following seeding onto prepared graft and culture surfaces was thought to correspond with the more rapid establishment of the flattened phenotype, as noted microscopically.16,33

Tissue Engineering of Prosthetic Vascular Grafts

Following flattening, where incomplete confluence results from seeding, the endothelial cells then migrate across the seeding surface to produce a monolayer. By some elegant work using removable discs, the migration of saphenous vein endothelial cells following establishment of healthy cultures was measured on a variety of culture surfaces.41 These showed that surface precoating not only improves initial adherence, but also promotes cell migration and development of confluence. Again, fibronectin was the most effective coating agent, with gelatin and extracellular matrix showing some advantage over uncoated culture flasks. Interestingly, laminin, despite being a constituent of the cell substratum in vivo, did not improve cell migration, implying that its role in the body is not primarily concerned with maintenance of endothelial monolayer integrity. In addition to enhancement of cell-surface adhesion, seeded cells also proliferate more rapidly on precoated surfaces,22 and some groups developed variations of the surface coatings specifically designed to improve graft endothelialization by promoting cell migration and proliferation.42 Coatings enriched with growth factors such as FGF have been shown to enhance ingrowth through the graft interstices43 and the development of an endothelial lining with a minimum of smooth muscle hyperplasia.44 The ability of small diameter vascular prostheses to adsorb and retain ECGF under flow conditions has been clearly demonstrated45 though in the rabbit model used, no differences in graft reendothelialization was seen at 1 month. Although the combination of growth factor-enhanced graft coatings and human endothelial cell seeding would appear to be a promising avenue of investigation, there have been no reports of such work.

Endothelial Cell Retention Under Flow Conditions The advantages of precoated over plain grafts also extend to dynamic seeding systems, in which the seeded grafts are exposed to the shear stress of flow, though different groups have achieved wide variation in results. A range of levels of flow, and thus shear stresses, have been used, including systems that approximate rates of flow seen in grafts.46,47 Electron microscopy of seeded grafts exposed to pulsatile flow systems show that a proportion of endothelial loss occurs in areas where the underlying graft is exposed by loss of surface coating,48 and that this is more marked with FN-coating than with preclot, which may account for the differences some workers have seen in levels of cell retention. The monolayer’s resistance to denudation under flow is a combination of the strength of the endothelial cell attachment to the matrix, and of the matrix to the graft. With regard to the former, factors previously observed to improve cell adherence such as duration of incubation, and also degree of confluence and interval prior to commencement of flow, are also relevant to resisting denudation under flow. Established endothelial monolayers formed by culture on the graft resist flow better than acutely attached cells,46,49,50 although previously noted problems with duration of incubation or graft preculture apply in terms of clinical applicability. Cell loss appears to occur in a biphasic pattern,47,51

Surface Precoating in the 1980s: A First Taste of Cell-Matrix Interactions

with an initial rapid rate presumably corresponding to loosely attached cells or areas of coating, which settles after a few minutes to a slower, more steady rate of loss (Fig. 4.4). However, we have noted that within a flow system sudden changes in the level of pulsatile shear stress result in a further period of rapid loss which then returns to the slower rate. The bond between coating and graft has been less thoroughly investigated, but seems to depend on the nature of the graft, with gelatin-coated Dacron performing significantly less well than ePTFE.38,40 Whether the FN-gelatin or the gelatin-Dacron interface is the important one is as yet uncertain, but clinically may not be particularly important, as ePTFE is the usual graft of choice in the infrainguinal region where it is expected that seeding will be most applicable. Up to 75% of applied matrix may be lost in the first 15 minutes after restoration of flow, but thereafter the graftmatrix bond appears stable within the time scale of the laboratory model. Electron microscopy of coated grafts after exposure to flow shows clearly that this early loss occurs at least partly by detachment of regions of matrix from the graft, presumably where the matrix-graft bond is less established, with matrix solubility probably contributing more towards the later, slow rate of matrix loss.

Seeding of Native Vascular Surfaces The difficulties in achieving an endothelial monolayer on prosthetic grafts immediately after seeding, given the almost inevitable shortfall in numbers of seeded cells, ethical and technical problems with cell or graft preculture, and the high losses from even FN-coated graft surfaces under flow, have prompted workers to examine the seeding potential of native arterial surfaces. The use of autologous cells and native conduits avoids any potential immune or foreign body reaction which might result in cell loss, and seeding following angioplasty or endarterectomy have both been considered. The advantages of seeding such regions are that the area for seeding is lower due to shorter segments being involved, thus increasing the seeding density without needing an increase in the numbers of seeded cells, and that cell at-

73

tachment and retention on autologous native surfaces might be expected to be optimal. Seeding after endothelial denudation by angioplasty balloon has been assessed in a rabbit iliac vein model52 using a double balloon catheter-based system and has shown that acute seeding with autologous cells on a relatively short time scale can restore a monolayer,53 which is maintained on restoration of flow. This monolayer retains functional integrity54 and may reduce neointimal hyperplasia.55 However, the relevance of these results to chronic arterial disease in humans remains uncertain, especially with previous experience of promising animal seeding studies. Endothelial cell seeding following endarterectomy in animals has also given hope, demonstrating most of the advantages seen in seeded prosthetic grafts.56-60 Work using human cell lines has demonstrated excellent cell attachment to both endarterectomy and vein angioplasty surfaces,61-63 and in the latter model flow studies have suggested cell retention kinetics similar to coated grafts. Following endarterectomy the exposed surface is very similar to that on which the endothelium normally rests and is high in FN, which probably accounts for the optimal attachment rates seen, and HAECs seeded onto endarterectomized vessels have been shown to be capable of forming a monolayer,61 again retaining functional integrity.64

Mechanism of Cell Adherence As a result of the improvements in seeding efficiency produced by surface coatings, increasing interest has developed in the interactions between endothelial cells and the matrix. The attachment and migration characteristics of the endothelial cell are due to the presence of a wide range of adhesion molecules and ligand-specific receptors on the cell membrane.65 Many of the receptors have multiple, overlapping roles in different pathways, or respond to more than one signal molecule. Variation in ligand affinity and patterns of receptor activation are common and the degree of complexity of the intracellular responses to the extracellular environment is only now becoming more fully

Fig. 4.4. Cell retention on seeded grafts under flow.

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appreciated. Some receptors are endothelial cell specific, some are shared with other vascular cells such as platelets or leukocytes, and some also occur on other vessel wall cell types including smooth muscle cells and fibroblasts. Because of their differential actions, different classes of adhesion molecules are frequently asymmetrically arranged between the luminal and abluminal cell membrane (Fig. 4.5). There are three main families of adhesion molecules present on endothelial cells, though classification is amended from time to time as new members are continually being identified and characterized, and many are known by more than one name. The three major adhesion molecules families consist of the selectins or leukocyte adhesion molecules (ELAMs), responsible for initial capture of circulating leukocytes in response to inflammatory mediators; the immunoglobulin superfamily, also implicated in cell-cell interactions; and the integrins, involved in both cell-substratum and cell-cell interactions. The selectins present on endothelial cells comprise two major groups; E-selectin or ELAM-1, and P-selectin or granule membrane protein-140 (GMP-140). These are induced by inflammatory mediators and promote the adherence of leukocytes to affected regions of endothelium via leukocyte surface ligands such as fucose groups. They are located on the luminal membrane surface and play little part in endothelium-substratum interactions. The immunoglobulin superfamily contains several classes, differing in the length and composition of a string of immunoglobulin domains which lie on the luminal side of the endothelial cell membrane, attached to a peptide chain that extends through this to a cytoplasmic terminal carboxyl group. They include the intercellular adhesion molecules (ICAMs) -1, -2, and -3; vascular cell adhesion molecule-1 (VCAM-1); and platelet endothelial cell adhesion molecule-1 (PECAM-1). ICAM-1 and VCAM-1 are inducible in response to circulating inflamma-

Fig. 4.5. Vascular endothelial cell adhesion molecules.

Tissue Engineering of Prosthetic Vascular Grafts

tory mediators, ICAM-2 and ICAM-3 are constitutively produced, but all bind to circulating leukocytes. PECAM-1 is constitutively produced and binds platelets in addition. Again, these molecules are principally involved in the endothelial response to inflammation, and these groups are therefore not further considered. Cell-substratum interactions such as the endothelial cell adherence to prosthetic grafts and coatings are mediated by the integrin family of adhesion molecules,66 which are 10- to 100-fold more common on the cell membrane than selectins or immunoglobulin-like adhesion molecules, and are present principally on the abluminal surface. Integrins are heterodimeric molecules, each subtype being composed of a noncovalently bonded pair of α and β subunits, each of which has multiple variants. There are several subclasses, distinguished by their β subunit, though their specific ligand is to a large extent determined by the α subunit. They bind with relatively low affinity to their ligands, made up for in large part by their frequency of expression, allowing multiple weak bonds to the substratum rather than a few strong links, and thus facilitating cell spread and migration without loss of firm adhesion, the ‘Velcro’ theory of cell attachment. Ligand binding is cation dependent, Ca2+ or Mg2+ depending on the integrin, bound to specific sites on the α subunit. The VLA group, identified by a β1 subunit, includes specific receptors found on the abluminal endothelial cell membrane for fibronectin, laminin and collagen. A β2 subunit identifies the leucam group which are found on circulating inflammatory cells and bind to ICAMs on endothelial cells. The cytoadhesin group of integrins, with a β3 subunit, includes low specificity, multiligand receptors binding fibrinogen, fibronectin, vitronectin, thrombospondin, laminin and vWF. Individual moieties with β4-β8 subunits binding a variety of basement membrane proteins, including those mentioned previously, have also been identified.

Surface Precoating in the 1980s: A First Taste of Cell-Matrix Interactions

The action of the integrins is a transmembrane link between the outside of the cell and the internal cytoskeleton. On ligand binding, the intracellular end of the β moiety binds to α actin, which stimulates assembly of attachment constructs, preventing the integrin from being pulled out of the cell. In addition to allowing motility and morphological changes by the endothelial cell, integrins can propagate extracellular matrix molecule orientation by the alignment of the resulting cytoskeletal elements, and by giving direction to the matrix molecules being laid down by the cell; this organization can be passed on to adjacent cells.

Nonendothelial Cell Lines Although most work has concentrated on the interactions between the seeding surface and the endothelial cells, the nature of the graft coating also has important effects on other cell types that come into contact. The ability of myofibroblasts and circulating macrophages to produce mitogens and growth factors that influence endothelial cells has been well documented,67 and these cell species can respond to stimulation by graft materials and coatings.68 This has been clearly shown in work on both polygalactin bioresorbable prostheses and Dacron, where macrophage activation from graft interaction results in release of FGF and the stimulation of endothelial ingrowth. Bioresorbable prostheses differ from Dacron in their ability to promote mitogenesis in the macrophages in addition to cytokine release,69 indicating differences between graft materials.

Future Directions Results from work on endothelial cell responses to stimuli have shown that the monolayer is subject, and responds, to a wide range of physical and biological factors. Both expansile stress from arterial pulsatility70,71 and the shear stress of flow46,72 elicit functional changes in the expression of biologically important factors by the endothelium. The cells are also subject to circulating agents, and locally produced paracrine factors deriving from leukocytes, platelets and macrophages.73-78 In addition, the nature of the underlying layer on which the monolayer is developed affects the endothelial cell function, a feature of both apparently passive basement membrane components,79-81 as well as other cell types in the supporting layers.82,83 Where the 1970s were the decade of developing viable cell culture, and the 1980s were primarily concerned with the kinetics of seeded cells’ interactions with the graft surface, the 1990s have so far shown an explosion in understanding of the molecular biology of the adhesion mechanisms of the endothelial cell and of the responses and relationships between viable seeded cells and the environment within which they are placed. Powerful research laboratory tools such as the polymerase chain reaction and advances in genetic manipulation techniques are proving important in the growing appreciation of the complexity of the largest organ in the body, and making sense of the active biological role of what was once thought to be a passive nonthrombogenic layer.

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References 1. Claggett GP, Burkel WE, Sharefkin JB et al. Platelet reactivity in vivo in dogs with arterial prostheses seeded with endothelial cells. Circ 1984; 69:632-636. 2. Ramberg K, Keough E, Callow AD et al. Indium-111 labeled platelet imaging of endothelial-cell-seeded smallcalibre synthetic grafts in the baboon. ASAIO J 1985; 8:95-98. 3. Birinyi LK, Douville EC, Lewis SA et al. Increased resistance to bacteraemic graft infection after endothelial cell seeding. J Vasc Surg 1987; 5:193-197. 4. Sharefkin JB, Latker C, Smith M et al. Early normalisation of platelet survival by endothelial cell seeding of dacron arterial prostheses in dogs. Surgery 1982; 92:385-393. 5. Budd JS, Allen KE, Hartley G et al. The effect of preformed confluent endothelial cell monolayers on the patency and thrombogenicity of small calibre vascular grafts. Eur J Vasc Surg 1991; 5:397-405. 6. Hess F, Steeghs S, Jerusalem R et al. Patency and morphology of fibrous polyurethane vascular prostheses implanted in the femoral artery of dogs after seeding with subcultivated endothelial cells. Eur J Vasc Surg 1993; 7:402-408. 7. Jarrell BE, Levine E, Shapiro S et al. Human adult endothelial cell growth in culture. J Vasc Surg 1984; 1:757-764. 8. Welch M, Durrans D, Carr HMH et al. Endothelial cell seeding; a review. Ann Vasc Surg 1992; 6(4):473-484. 9. Sharefkin JB, Lather C, Smith M et al. Endothelial cell labelling with indium-111 oxine as a marker of cell attachment to bioprosthetic surfaces. J Biomed Mater Res 1983; 17:345-357. 10. Vohra R, Thomson GJL, Carr HMH et al. Comparison of different vascular prostheses and matrices in relation to endothelial cell seeding. Br J Surg 1991; 78:417-420. 11. Williams SK, Jarrell BE, Friend L et al. Adult human endothelial cell compatibility with prosthetic graft material. J Surg Res 1985; 38:618-629. 12. Campbell JB, Lundgren C, Herring MB et al. Attachment and retention of Indium-111-labeled endothelial cells onto polyester elastomer. ASAIO J 1985; 8:113-117. 13. Radomski JS, Jarrell BE, Williams SK et al. Initial adherence of human capillary endothelial cells to Dacron. J Surg Res 1987; 42:133-140. 14. Budd JS, Allen KE, Bell PRF. Effects of two methods of endothelial cell seeding on cell retention during blood flow. Br J Surg 1991; 78:878-882. 15. Kent KC, Oshima A, Whittemore AD. Optimal seeding conditions for human endothelial cells. Ann Vasc Surg 1992; 6(3):258-264. 16. Jarrell BE, Williams SK, Solomon L et al. Use of an endothelial monolayer on a vascular graft prior to implantation: Temporal dynamics and compatibility with the operating room. Ann Surg 1986; 203:671-678. 17. Tannenbaum G, Ahlborn T, Benvenisty A et al. High density seeding of cultured endothelial cells leads to rapid coverage of polytetrafluoroethylene grafts. Curr Surg 1987:318-321. 18. Watkins MT, Sharefkin JB, Zajtchuk R et al. Adult human saphenous vein endothelial cells: Assessment of their reproductive capacity for use in endothelial seeding of vascular prostheses. J Surg Res 1984; 36:588-596. 19. Jarrell BE, Williams SK, Stokes G et al. Use of freshly isolated capillary endothelial cells for the immediate establishment of a monolayer on a vascular graft at surgery. Surgery 1986; 100:392-399.

76 20. Nicholson LJ, Clarke JM, Pittilo RM et al. The mesothelial cell as a non-thrombogenic flow surface. Thromb Haemost 1984; 52(2):102-104. 21. Ramalanjoana G, Kempczinski RF, Rosenman JE et al. The effect of fibronectin coating on endothelial cell kinetics in polytetrafluoroethylene grafts. J Vasc Surg 1986; 3:264-272. 22. Sank A, Rostami K, Weaver F et al. New evidence and new hope concerning endothelial seeding of vascular grafts. Am J Surg 1992; 164:199-204. 23. Stansby G, Shukla N, Fuller B et al. Seeding of human microvascular endothelial cells onto PTFE graft material. Br J Surg 1991; 78:1189-1192. 24. Thomson GJL, Vohra R, Walker MG. Cell seeding for small diameter ePTFE grafts: A comparison between adult human endothelial and mesothelial cells. Ann Vasc Surg 1989; 3(2):140-145. 25. Vohra R, Thomson GJL, Carr HMH et al. In vitro adherence and kinetics studies of adult human endothelial cell seeded polytetrafluoroethylene and gelatin-impregnated Dacron grafts. Eur J Vasc Surg 1991; 5:93-103. 26. Thomson GJL, Vohra RK, Carr HMH et al. Adult human endothelial cell seeding using expanded polytetrafluoroethylene vascular grafts: A comparison of four substrates. Surgery 1991; 109:20-27. 27. Pratt KJ, Jarrell BE, Williams SK et al. Kinetics of endothelial cell surface attachment forces. J Vasc Surg 1988; 7:591-599. 28. van Wachem PB, Beugeling T, Mallens BW et al. Deposition of endothelial fibronectin on polymeric surfaces. Biomaterials 1988; 9:121-123. 29. van Wachem PB, Mallens BW, Dekker A et al. Adsorption of fibronectin derived from serum and from human endothelial cells on to tissue culture polystyrene. J Biomed Mater Res 1987; 21:1317-1327. 30. Seeger JM, Klingman N. Improved endothelial cell seeding with cultured cells and fibronectin-coated grafts. J Surg Res 1985; 38:641-647. 31. Kesler KA, Herring MB, Arnold MP et al. Enhanced strength of endothelial attachment on polyester elastomer and polytetrafluoroethylene graft surfaces with fibronectin substrate. J Vasc Surg 1986; 3:58-64. 32. Curti T, Pasquinelli G, Preda P et al. An ultrastructural and immunocytochemical analysis of human endothelial cell adhesion on coated vascular grafts. Ann Vasc Surg 1989; 3(4):351-363. 33. Kaehler J, Zilla P, Fasol R et al. Precoating substrate and surface configuration determine adherence and spreading of seeded endothelial cells on PTFE grafts. J Vasc Surg 1989; 9:535-541. 34. Baker KS, Williams SK, Jarrell BE et al. Endothelialization of human collagen surfaces with human adult endothelial cells. Am J Surg 1985; 150:197-200. 35. Dalsing MC, Kevorkian M, Raper B et al. An experimental collagen-impregnated Dacron graft: Potential for endothelial seeding. Ann Vasc Surg 1989; 3(2):127-133. 36. Zilla P, Fasol R, Preiss P et al. Use of fibrin glue as a substrate for in vitro endothelialization of PTFE vascular grafts. Surgery 1989; 105:515-522. 37. Mosquera DA, Goldman M. Endothelial cell seeding. Br J Surg 1991; 78:656-660. 38. Carr HMH, Vohra R, Welch M et al. Fibronectin binding to gelatin-impregnated Dacron (Gelseal) prostheses. Artif Organs 1992; 16:342-345.

Tissue Engineering of Prosthetic Vascular Grafts 39. Budd JS, Allen KE, Bell PRF et al. The effect of varying fibronectin concentration on the attachment of endothelial cells to polytetrafluoroethylene grafts. J Vasc Surg 1990; 12:126-130. 40. Vohra RK, Thompson GJL, Sharma H et al. Fibronectin coating of expanded polytetrafluoroethylene (ePTFE) grafts and its rôle in endothelial seeding. Artif Organs 1990; 14:41-45. 41. Hasson JE, Wiebe DH, Sharefkin JB et al. Migration of adult human vascular endothelial cells: Effect of extracellular matrix proteins. Surgery 1986; 100:384-390. 42. Lalka SG, Oelker LM, Malone JM et al. Acellular vascular matrix: A natural endothelial cell substrate. Ann Vasc Surg 1989; 3(2):108-117. 43. Greisler HP, Cziperle DJ, Kim DU et al. Enhanced endothelialization of ePTFE grafts by fibroblast growth factor type 1 pretreatment. Surgery 1992; 112:244-254. 44. Gray JL, Kang SS, Zenni GC et al. FGF-1 affixation stimulates ePTFE endothelialization without intimal hyperplasia. J Surg Res 1994; 57:596-612. 45. Greisler HP, Klosak JJ, Dennis JW et al. Biomaterial pretreatment with ECGF to augment endothelial cell proliferation. J Vasc Surg 1987; 5:393-399. 46. Sentissi JM, Ramberg K, O’Donnell TFJ et al. The effect of flow on vascular endothelial cells grown in tissue culture on polytetrafluoroethylene grafts. Surgery 1986; 99:337-342. 47. Vohra RK, Thomson GJL, Sharma H et al. Effects of shear stress on endothelial cell monolayers on expanded polytetrafluoroethylene (ePTFE) grafts using preclot and fibronectin matrices. Eur J Vasc Surg 1990; 4:33-41. 48. James NL, Schindhelm K, Slowiaczek P et al. Endothelial cell seeding of small diameter vascular grafts. Artif Organs 1990; 14:355-360. 49. Miyata T, Conte MS, Trudell LA et al. Delayed exposure to pulsatile shear stress improves retention of human saphenous vein endothelial cells on seeded ePTFE grafts. J Surg Res 1991; 50:485-493. 50. Schneider PA, Hanson SR, Price TM et al. Durability of confluent endothelial cell monolayers on small calibre vascular prostheses in vitro. Surgery 1988; 103:456-462. 51. Rosenman JE, Kempczinski RF, Pearce WH et al. Kinetics of endothelial cell seeding. J Vasc Surg 1985; 2:778-784. 52. Thompson MM, Budd JS, Eady SL et al. A method to transluminally seed angioplasty sites with endothelial cells using a double balloon catheter. Eur J Vasc Surg 1993; 7:113-121. 53. Thompson MM, Budd JS, Eady SL et al. Platelet deposition after angioplasty is abolished by restoration of the endothelial cell monolayer. J Vasc Surg 1994; 19:478-486. 54. Thompson MM, Budd JS, Eady SL et al. Endothelial cell seeding of damaged native vascular surfaces: Prostacyclin production. Eur J Vasc Surg 1992; 6:487-493. 55. Thompson MM, Budd JS, Eady SL et al. The effect of transluminal endothelial seeding on myointimal hyperplasia following angioplasty. Eur J Vasc Surg 1994; 8:423-434. 56. Bush HL, Jakubowski JA, Curl GR et al. Luminal healing of arterial endarterectomy: Role of autogenous endothelial cell seeding. Surg Forum 1985; 36:446-450. 57. Bush HL, Jakubowski JA, Sentissi JM et al. Neointimal hyperplasia occurring after carotid endarterectomy in a canine model: Effect of endothelial cell seeding vs perioperative aspirin. J Vasc Surg 1987; 5:118-125. 58. Schneider PA, Hanson SR, Price TM et al. Confluent durable endothelialisation of endarterectomised baboon aorta

Surface Precoating in the 1980s: A First Taste of Cell-Matrix Interactions by early attachment of cultured endothelial cells. J Vasc Surg 1990; 11:365-372. 59. Krupski WC, Bass A, Anderson JS et al. Aspirin-independent antithrombotic effects of acutely attached cultured endothelial cells on endarterectomised arteries. Surgery 1990; 108:283-291. 60. Sterpetti AV, Schultz RD, Bailey RT. Endothelial cell seeding after carotid endarterectomy in a canine model reduces platelet uptake. Eur J Vasc Surg 1992; 6:390-394. 61. Smyth JV, Rooney OB, Dodd PDF et al. Culture of human endothelial cells on endarterectomy surfaces. Eur J Vasc Endovasc Surg 1995; 10:308-315. 62. Walluscheck KP, Steinhoff G, Haverich A. Endothelial cell seeding of native vascular surfaces. Eur J Vasc Endovasc Surg 1996; 11:290-303. 63. Thompson MM, Budd JS, Eady SL et al. Effect of seeding time and density on endothelial cell attachment to damaged vascular surfaces. Br J Surg 1993; 80:359-362. 64. Smyth JV, Rooney OB, Dodd PD et al. Production of plasminogen activator inhibitor-1 by human saphenous vein endothelial cells seeded onto endarterectomy specimens in vitro. J Vasc Surg 1997; 25:722-725. 65. Mousa SA, Cheresh DA. Recent advances in cell adhesion molecules and extracellular matrix proteins: Potential clinical implications. DDT 1997; 2:187-199. 66. Extracellular matrix receptors on animal cells: The integrins. In: Alberts B, Bray D, Lewis J et al. Molecular biology of the cell. 3rd ed. New York: Garland Publishing, Inc, 1994:995-1000. 67. Sterpetti AV, Crucina A, Napoli F et al. Growth factor release by smooth muscle cells is dependent on haemodynamic factors. Eur J Vasc Surg 1992; 6:636-638. 68. Greisler HP, Henderson SC, Lam TM. Basic fibroblast growth factor production in vitro by macrophages exposed to Dacron and polyglactin 910. J Biomater Sci Polym Ed 1993; 4:415-430. 69. Greisler HP, Dennis JW, Endean ED et al. Macrophage/ biomaterial interactions: The stimulation of endothelialization. J Vasc Surg 1989; 9:588-593. 70. Sumpio BE, Widmann MD, Ricotta J et al. Increased ambient pressure stimulates proliferation and morphological changes in cultured endothelial cells. J Cell Physiol 1994; 158:133-139. 71. Iba T, Sumpio BE. Tissue plasminogen activator expression in endothelial cells exposed to cyclic strain in vivo. Cell Transpl 1992; 1:43-50.

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72. Eskin SG, Ives CL, Frangos JA et al. Cultured endothelium: The response to flow. ASAIO J 1985; 8:109-112. 73. Kugiyama K, Sakamoto T, Misumi I et al. Transferable lipids in oxidized low-density lipoprotein stimulate plasminogen activator inhibitor-1 and inhibit tissue-type plasminogen activator release from endothelial cells. Circ Res 1993; 73:335-343. 74. Yamamoto C, Kaji T, Sakamoto M et al. Effect of endothelin on the release of tissue plasminogen activator and plasminogen activator inhibitor-1 from cultured human endothelial cells and interaction with thrombin. Thromb Res 1992; 67:619-624. 75. Loskutoff DJ, Ny T, Sawdey M et al. Fibrinolytic system of cultured endothelial cells; regulation by plasminogen activator inhibitor. J Cell Biochem 1986; 32:273-280. 76. van Hinsbergh VWM, Bertina RM, van Wijngaarden A et al. Activated protein c decreases plasminogen activator inhibitor activity in endothelial cell-conditioned medium. Blood 1985; 65:444-451. 77. van Hinsbergh VWM, Kooistra T, van den Berg EA et al. Tumor necrosis factor increases the production of plasminogen activator inhibitor in human endothelial cells in vitro and in rats in vivo. Blood 1988; 72:1467-1473. 78. Hanss M, Collen D. Secretion of tissue-type plasminogen activator and plasminogen activator inhibitor by cultured human endothelial cells: Modulation by thrombin, endotoxin and histamine. J Lab Clin Med 1987; 109:97-104. 79. Lundgren CH, Herring MB, Arnold MP et al. Fluid shear disruption of cultured endothelium: The effect of cell species, fibronectin crosslinking and supporting polymer. ASAIO Trans 1986; 32:334-338. 80. Ives CL, Eskin SG, McIntire LV et al. The importance of cell origin and substrate in the kinetics of endothelial cell alignment in response to steady flow. ASAIO Trans 1983; 29:269-274. 81. Li JM, Menconi MJ, Wheeler HB et al. Precoating ePTFE grafts alters production of endothelial cell derived thrombomodulators. J Vasc Surg 1992; 15:1010-1017. 82. Christ G, Seiffert D, Hufnagl P et al. Type 1 plasminogen activator inhibitor synthesis of endothelial cells is downregulated by smooth muscle cells. Blood 1993; 81:1277-1283. 83. Scott-Burden T, Vanhoutte PM. Regulation of smooth muscle cell growth by endothelium-derived growth factors. Tex Heart Inst J 1994; 21:91-97.

CHAPTER 5 Surface Precoating in the 1990s: The Fine Tuning of Endothelial Cell Transplantation Mark M. Samet, Victor V. Nikolaychik, Peter I. Lelkes

Introduction

E

ach year 900,000 people in the USA alone suffer from arterial disorders requiring some form of surgical intervention. Over half of these procedures involve peripheral reconstruction using vascular grafts to bypass an occluded vessel segment. However, 60% of the patients undergoing vascular surgery do not have a suitable vessel for grafting and, therefore, need synthetic grafts.1 In this realm, about 350,000 artificial vascular grafts are implanted annually, and the numbers are expected to rise rapidly in the future. In spite of the increasing demand for vascular prostheses, the long term success of their use as arterial substitutes is, at best, mixed.2,3 In artery bypasses, with bore sizes ≥ 5 mm in diameter Dacron® and ePTFE grafts demonstrate cumulative patency rates similar to those of the natural grafts.1 But, as small artery substitutes, these prostheses have performed poorly and, to date, no vascular prosthesis of ≤ 4 mm in diameter remains patent for more than a couple of months.1 The high failure rates of small diameter prosthetic grafts are believed to be intimately associated with the poor healing and rejection process of the implanted conduit.3-7 One of the key determinants in this “cause and effect” relationship is the problem of insufficient hemocompatibility of the polymeric graft material.8 Insufficient hemocompatibility of the biomaterials is also a pressing issue in the field of cardiac prostheses. Each year, end-stage heart diseases claim some 75,000 victims in the USA alone. Although 28,000 of these patients qualify for heart transplants, the supply of donated hearts is limited to approximately 2,000 hearts per year.9,10 Thus, the majority of these patients has no realistic hope of long term survival, unless permanent total artificial hearts (TAH) or ventricular assist devices (VAD) become available. The vision of developing a permanent cardiac prosthesis is more than a half century old. The first practical implementation of this idea was achieved in the 1950s by W.J. Kolff and his collaborators, who kept a dog alive for 90 min after successfully replacing the native heart with an artificial device.11 By the mid-1970s, the first generation of TAHs and VADs was undergoing advanced clinical trials.11-14 These units were activated externally by pneumatic drivers and connected trough transcutaneous pressure lines. Besides restricting the patient’s postoperative quality of life, these lines served as a major source for bacterial infection and initiated thromboembolic complications. The newest generation of cardiac prostheses, designed to be totally implantable, are powered by a compact electrical source.14 In spite of significant mechanical and electronic improvements in the new designs, the permanent or even temporary clinical use of cardiac prostheses as a bridge to recovery15 is still Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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Tissue Engineering of Prosthetic Vascular Grafts

thwarted by serious obstacles such as infections and thromboembolic complications which are manifested in stroke and multiple organ failures.16,17 As in the case of artificial vascular grafts, many of the complications with cardiac prostheses can be traced back to the problem of inadequate hemocompatibility of the blood-contacting biomaterials in these devices. Thus, it follows that if we can successfully overcome the blood compatibility barrier, we shall markedly improve the clinical prospects of both artificial hearts and small diameter prosthetic grafts.

lium by mechanical forces.38,39 Many investigators, however, turned to surface treatment with select extracellular matrix (ECM) proteins or blood plasma proteins (“fibrin glue”), with or without supplements, as a means for improving cellular performance in grafts.40,41 Our ongoing studies focus on surface coating as a means for optimizing endothelialization.42,43 Specifically, our motivation is to form a multifunctional protein complex that will perform as a biointegrator between the living cells and the underlying polymeric scaffold in cardiac prostheses.

Endothelialization: Past and Current Approaches

Fine Tuning of Endothelialization via Surface Coating

The interaction of blood with “medical-grade” biomaterials commonly results in thrombus formation and/or thromboembolic complications. After decades of disappointing attempts to passivate the synthetic blood-contacting surfaces by chemical modifications or by generation of a pseudoneointima,15,18 the focus has turned towards establishing on the synthetic blood-contacting surfaces “nature’s own blood-compatible container”,19 i.e., reconstructing the endothelial cell (EC) lining of the vascular wall. The rationale for covering the prosthetic luminal surface with a viable layer of ECs, i.e., endothelialization, is rather simple: in vivo, ECs present under normal physiologic conditions an anti-thrombogenic surface which is pivotal to the prevention of fibrin deposition and inhibition of platelet adhesion on the vessel wall. Thus, it has been reasoned that endothelialization of cardiovascular prostheses would reduce the thrombogenicity of these devices. In addition, adverse immunogenic responses from the host would be avoided by transplanting autologous cells. Indeed, these principles were successfully demonstrated more than two decades ago in artificial grafts.20 However, in 1975, the art of procuring and culturing endothelial cells was in its infancy, and very little was known about EC-biomaterial compatibility. Since then, together with significant advances in endothelial cell biology, molecular biology, and biotechnology,21,22 a fairly thorough understanding has been gained of the basic requirements for successfully growing endothelial cells on the biomaterials used in cardiovascular prostheses: PTFE, ePTFE, Dacron®, and various brands of polyurethanes (PUs).23-26 In recent years, endothelialization of vascular graft surfaces with autologous ECs may have solved several important problems in cardiovascular surgery27-31 and paved the way for the development of new areas, such as biohybrid organs32 and gene therapy.33-36 Thus, at present, the question is no longer whether or not endothelialization is feasible, but, rather, how it can be optimized. Optimization, or fine tuning of endothelialization is a multifaceted task. It involves, but is not limited to, improvement of endothelial cell attachment and growth on the synthetic surface, enhancement of cell adaptation to and subsequent endurance in a hemodynamic environment, and finally, promotion of the anti-coagulant repertoire of the entire EC lining. Several researchers have used genetically modified ECs as a vehicle for achieving these goals,23,33 while others focused on increasing the hydrophilicity of the synthetic scaffold37 or preconditioning the cultured endothe-

The best biointegrating interface for ECs is their native basement membrane which, in vivo, contributes to the maintenance of the architectural and functional integrity of the endothelium.44 This basement membrane is comprised of structural proteins such as collagens, as well as adhesive proteins such as laminin (LM) and fibronectin (FN). These and several other ECM proteins regulate, via RGD (arginine-glycine-aspartic acid)-based integrin-mediated reception, cell adhesion and spreading.45 In addition, many of these ECM molecules are also actively synthesized during cell proliferation/migration.46 Using ECM-contained adhesive proteins, various researchers have employed FN, LM, or collagen iv (Col IV), or a combination thereof, as a surface coating to facilitate endothelialization of biomaterials.40,47,48 Similarly, short RGD peptides, or their mimics, have been covalently coupled to the PU backbone to achieve the same goal.49 We revisited the effectiveness of select ECM-contained adhesive monoproteins to promote cells growth on PUs. For this purpose we seeded bovine aortic ECs (BAEC) at a density of 25,000 cells/cm2 onto PU substrates and monitored their growth for several days. As a control, we cultured BAECs on tissue culture grade polystyrene (TC) in 24 well plates (Costar, Cambridge, MA). The PU samples were made of Chronoflex, and fitted into the wells of 24 well plates as previously described.50 The surface of the PUs was either left uncoated or was coated by passive adsorption for 1 h with FN (20 µg/ml, pH=7.4) or Col IV (300 µg/ml, pH=3.0) prior to cell seeding. The results of this study are summarized in Figure 5.1a. Each data point (mean ± standard deviation) represents the number of cells per cm2 from 3 independent experiments (6 wells per case). As expected, the cells grew poorly on plain (uncoated) PU, and their number by day nine was significantly lower than in controls (0.92 ± 0.1 x 105 vs. 1.96 ± 0.8 x 105, p < 0.05). Proliferation of BAEC improved significantly upon coating the PUs with either FN or Col IV, but it did not match the rate of growth on TC. Interestingly, there was no substantial difference in the rate of growth between cells cultured on FN or Col IV coatings. In fact, on day 9, the number of cells in both cases was almost equal (1.77 ± 0.87 x 105 on FN vs. 1.69 ± 0.11 x 105 on Col IV, p=0.425). Thus, our data confirm the findings of others that biomaterials coated with ECM proteins are more cytocompatible than their untreated counterparts.40,51,52 Moreover, our results also imply that in coating the polymeric scaffold with only one kind of ECM adhesive protein,

Surface Precoating in the 1990s: The Fine Tuning of Endothelial Cell Transplantation

the particular type of this protein is of secondary importance for ECs grown in a static environment. Subsequently, we examined whether the effectiveness of a particular surface coating in promoting cell growth may depend on the method of protein immobilization onto the polymeric surface. To this end, we seeded BAEC at a density of 25,000 cells/cm2 onto PUs coated by different means with ECM proteins, and monitored their growth for several days. The PU samples consisted of 4 separate groups: Group #1: PU coated with FN by adsorption. The starting concentration of FN was 20 µg/ml (pH=7.4). After 1 h of incubation at 37°C, the final concentration on the surface, as determined by the sulforhodamine B assay,53 was 20 ng/cm2. Group #2: PU coated with FN by photoimmobilization using the method of Clapper et al.54 The final concentration of FN on the surface was as in group #1. Group #3: PU coated with Col IV by adsorption. The starting concentration was 300 µg/ml at pH = 3.0. After 1 h of incubation at 37°C, the final concentration on the surface was 50 ng/cm2. Group #4: PU coated with Col IV by photoimmobilization. The final concentration of Col IV on the surface was as in group #3. As shown in Figure 5.1b, the cells grew equally well on FN substrates, whether adsorbed or photoimmobilized. On the other hand, when Col IV-based substrates were used, cell growth was significantly better on PU coated by adsorption than by photoimmobilization (1.69 ± 0.11 x 105 vs. 1.33 ± 0.11 x 105, p < 0.05). Thus, for select proteins some methods of immobilization may indeed result in better substrates for cell growth than others. However, in our hands, no ECMderived, monoprotein coating of PU can match the conditions offered by a tissue culture-grade polystyrene. Clearly, a substrate comprised of two dimensionally assembled, ECM adhesive monoproteins does not provide for an adequate surface coating that can be used as a vehicle for optimizing endothelialization of cardiovascular prostheses. Hence, more sophisticated approaches are required to generate a

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biointegrated proteinaceous meshwork at the PU surface which would structurally mimic the topology of the subendothelial basement membrane and also provide space for the incorporation of the endogenous, EC-derived extracellular proteins. Such a provisional ECM meshwork might be found in fibrin glue, the reaction product of fibrinogen and thrombin. In the past, fibrin glue was shown to be beneficial in supporting both EC attachment and growth on synthetic materials40,42,43,55-57 and for establishing a stable, shear-resistant monolayer on ePTFE grafts.58 To improve this coating we substituted purified fibrinogen with plasma cryoprecipitate. Indeed, the results of those studies were encouraging: After 4 days in culture the number of cells on fibrinbased coating was higher by almost 45% than on TC. Cell coverage was even higher when the fibrin-based coating was supplemented with heparin and insulin.42 Based on these findings, we assessed the efficacy of the fibrin-based coating to promote EC growth on the luminal surface of PU-covered “ersatz” ventricles, (T-25 tissue culture flasks61). For comparison purposes we also used “ersatz” ventricles covered with either plain or FN-coated PU. In all cases the luminal surface of the PUs used (Biospan™)50 was smooth and the coatings were prepared without supplements such as growth factors, heparin, or insulin. The study was conducted in two sequential steps to realistically mimic a full scale endothelialization of the blood sac of a VAD. First, the cells (BAECs) were seeded with rotation for 3 h onto either precoated or plain PU surfaces.61 Thereafter, the “ersatz” ventricles were removed from the rotator and kept in an upright position, and the adherent ECs were allowed to spread and proliferate for several days. We monitored cell growth every other day by Alamar Blue assay.50 The results obtained are presented in Figure 5.2a. As expected, the lowest rate of growth was obtained on plain PU. On FN-coated PU the cells grew somewhat better, presumably because there were more cells attached immediately after seeding and so they could interact more effectively with one another. In the past, several studies had already shown that such interactions play an important role

Fig. 5.1. Cell growth on various substrates under static conditions: (a) Effect of select surface treatments on cell proliferation. FN or Col IV were immobilized onto PU samples by adsorption as detailed in the text. (b) Effect of protein immobilization methods on cell densities. FN or Col IV were immobilized onto PU samples either by adsorption or photoimmobilization as detailed in the text. Cell number was quantitated every other day by Alamar Blue assay.

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in cell proliferation in vitro.62,63 By contrast, cells seeded on plasma cryoprecipitate-made fibrin substrate (CRY) grew rapidly and reached confluence by the seventh day in culture. Importantly, purified fibrinogen-made fibrin substrate (FBG) was not as effective as CRY in promoting cell growth; even on the ninth day in culture the number of cells on FBG was still significantly lower than on CRY (2.86 ± 0.28 x 105 vs. 3.41 ± 0.25 x 105, p < 0.05). Since both coatings are comprised of a fibrin-based matrix, it is likely that some of the intrinsic plasma proteins contained in CRY may have contributed to the enhanced rate of cell growth on this substrate. Smooth-surface PUs have been used in cardiac prostheses as a means of preventing platelet aggregation and thrombus formation on the blood-contacting surfaces. However, notwithstanding the improvements in PU manufacturing and surface processing, the microscopic appearance of the luminal surface in these devices has never been completely smooth, and also the hemocompatibility of the smooth surfaces has not met the initial expectations. On the other hand, textured PU surfaces have been advocated as a scaffold of choice for enhancing pseudoneointima formation64,65 and for promoting endothelial cell attachment onto the luminal surface of cardiovascular devices.29,66,67 In adopting the latter concept for the endothelialization of VAD blood sacs, we evaluated CRY enhancement of cell growth on textured PU. The experiments were conducted in the same manner as above with smooth, uncoated PUs serving as controls. Figure 5.2b exemplifies our findings for plain and CRYcoated PUs after seven days in culture. BAECs proliferated much faster on the uncoated, textured PU than on its smooth counterpart, (2.17 ± 0.21 x 105 vs. 1.49 ± 0.21 x 105, p < 0.05). Again, the significant difference in the number of cells covering the surfaces may have resulted, in part, from more efficient attachment of cells on textured surface than on smooth PU. Application of CRY coating masked the topological difference between the two types of surfaces and enhanced cell growth with the same effectiveness. Thus, the number of cells on both PUs was virtually the same at all times; for example, on the seventh day in culture (Fig. 5.2b) we measured 3.24 ± 0.23 x 105 cells/cm2 on smooth PU+CRY vs. 3.56 ± 0.23 x 10 5 cells/cm2 on textured PU+CRY Fig. 5.2. Cell growth on various substrates in PU-covered “ersatz” ventricles after seeding with rotation: (a) Effect of select surface coatings on cell proliferation. (b) Cell densities in cultures grown on smooth and textured PUs, with or without CRY coating. “ersatz” ventricles, consisting of PUcovered T-25 flasks, were seeded with rotation for 3 h at an inoculation density of 60,000 cells/cm2. After stopping the seeding, the adherent BAECs were allowed to spread and proliferate under static conditions. Cell number was assessed every other day by Alamar Blue assay, and the cultures were refed after completion of the assay.

Tissue Engineering of Prosthetic Vascular Grafts

(p = 0.24). On both CRY-coated PUs, the number of cells was about 1.5 times higher than on plain, textured PUs. These data suggest that fibrin-based coating is indeed a preferred substrate for optimizing EC growth on the luminal surface of cardiovascular prostheses. In addition to enhancing cell growth, increasing the initial efficiency of cell attachment onto biomaterials is pivotal for successfully establishing an EC monolayer in prosthetic devices. In the past, we and others have explored various ways for increasing the affinity of synthetic surfaces to cells, for example, by glow discharge treatment,68 coating with substrates composed of select proteins,40,69,70 or both.61 Indeed, when BAECs were seeded under static conditions onto various PU substrates, we saw a substantial improvement in attachment efficiency on treated surfaces vs. plain PU (Fig. 5.3a). Our study also indicated that the most suitable surface for cell attachment was TC; CRY-coated PU was second to the best (94.1 ± 3.23%, p < 0.05). In our hands, the performance of CRY (and FBG) coatings was dependent upon the concentration of fibrinogen employed. As demonstrated in Table 5.1, cell attachment efficacy of purified fibrinogen varied in a bell-shaped manner between concentration levels of 0-30 mg/ml. The best performance was obtained at 5 mg/ml but, for preparing a coating of controllable thickness on a textured surface, we used only 3.5 mg/ml. Optimal adherence of endothelial cells under static conditions onto flat PU substrates or a TC surface does not necessarily translate into efficient attachment of cells during endothelialization under rotation of a complex geometric structure such as a cardiovascular prosthesis. In our hands, there are four equally important factors that determine the efficiency of cell attachment under these unique conditions: number of cells in the inoculum, speed of rotation, duration of seeding and surface treatment by coating. By fine tuning these parameters we were able to improve the homogeneity of surface coverage, optimize the time of residence/contact of cells with the substrate, and thus promote cell attachment. We established the following as optimal for lining our artificial ventricles: speed at 10 rotations/h, duration of seeding at 3 h, and inoculation density at 6.0 x 104 cells/cm2. Utilizing these settings, we examined the efficiency of attachment of cells that were seeded onto

Surface Precoating in the 1990s: The Fine Tuning of Endothelial Cell Transplantation

various substrates in smooth PU-covered “ersatz” ventricles. Plain TC served as controls. The results obtained are shown in Figure 5.3b. In contrast to the findings under static conditions, plain TC was the least suitable substrate for cell attachment under rotation, whereas CRY ranked the best (553.1 ± 21% over TC). FBG ranked second to the best (328.2 ± 13% over TC) and FN-based substrate was the third with a considerably lower efficiency (257.3 ± 11% over TC). The topography of the luminal surface in most vascular grafts is primarily rough.71,72 Also, some of the TAHs and VADs that are currently tested or developed, including the Heartmate® and the Milwaukee Heart™, have a rough luminal surface. To gain a more factual representation of the attachment efficiencies during endothelialization of artificial devices, we repeated the above experiments using “ersatz” ventricles covered with textured PU. Figure 5.4a presents a typical SEM photomicrograph of a textured surface that was prepared, as previously reported,43 by sprinkling NaCl crystals (in this case the mean diameter was 32 µm) over a wet PU cast. The entire surface appears corrugated by

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abundant, randomly dispersed cavities of varying dimensions, ranging in size from 30-120 µm in diameter. Figure 5.4b exemplifies attachment efficiencies on plain and CRYcoated PUs. It is apparent from panel A that cell attachment onto an uncoated, textured surface was significantly better than onto plain smooth PU (134.8 ± 8% and 108.2 ± 6%, respectively, compared to TC controls, p < 0.05). This improvement was likely due to the increase in cell residence time at the site of contact and isolation from fluid currents, rather than resulting from enhancement in surface affinity. In the case of CRY-coated PUs, the fibrin meshwork completely filled the pores in the textured PU and masked the topological differences between rough and smooth surfaces (Figs. 5.4 c,d). The resulting substrates had a finely patterned luminal surface that was enriched in numerous RGD binding sites inherent to the component proteins. As shown in panel B of Figure 5.4b, cells adhered readily onto both these substrates, leading to similar attachment efficiencies of more than 500% above control level. Thus, by using CRY coating, we completely eliminated the difference in attachment efficiencies between smooth and rough PUs. Having optimized cell attachment and growth on PUs, we determined whether the CRY coating could also enhance the retention of ECs in a monolayer that is exposed to dynamic conditions. For these in vitro studies, we used VAD blood sacs which were fabricated from a commercially available PU (Biospan™) by the lost wax technique.43 For cell seeding, the lumina of fully assembled VADs were filled with 125 ml of Air/CO2-presaturated culture medium containing 4.0 x 107 BAECs, mounted onto the rotating arm of the seeding apparatus61 and rotated at 10 rotations/h for 3 h. Following seeding, the VADs were placed, in an upright position, in a standard CO2 incubator. Cell proliferation was measured every other day, and the cultures were refed after conclusion of each Alamar Blue assay.50 Two days after seeding, the monolayers in the blood sacs had reached confluence. We then maintained them for eight additional

Table 5.1 Efficiency of cell attachment as a function of fibrinogen concentration Concentration of Fibrinogen (mg/ml)

Fig. 5.3. Efficiency of cell attachment onto various substrates: (a) under static conditions, and (b) after seeding with rotation of “ersatz” ventricles. BAECs were inoculated at a density of 60,000 cells/cm2. After 3 h of seeding, the cultures were washed gently with medium to remove nonadherent cells, and assayed by Alamar Blue assay to determine the number of cells attached to the surface. The resulting numbers were normalized to data obtained on TC surface.

0.0 0.5 1.0 2.0 5.0 10.0 15.0 20.0 30.0

Efficiency of Cell Attachment (%) a) 61.48 ± 72.13 ± 76.28 ± 83.56 ± 84.01 ± 83.72 ± 79.56 ± 76.44 ± 72.11 ±

4.87 1.87 3.14 3.14 2.91 3.47 2.43 2.05 1.52

aThe study was conducted under static conditions and the number

of cells recorded at each fibrinogen concentration was normalized to the number of cells obtained on plain TC. The fibrinogen used in this particular study was purified in the Hemostasis Research Laboratory at Sinai Samaritan Medical Center, Milwaukee, WI, and was kindly donated by its director, Dr. M. Mosesson.

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days under static conditions to allow the cells to synthesize, assemble, and mature their ECM on the underlying scaffold.73 Following a total of eleven days in a static environment, the monolayers were exposed for 24 h to one of the following dynamic conditions: 1. Flexing of the blood sac wall at 1 Hz between systole and diastole positions, with minimal flow of medium inside the blood sac; 2. Perfusion of the blood sac in an assembled VAD with pulsatile flow at the mean rate of 500 ml/min (blood sac wall was maintained motionless during the whole period); and 3. Pumping action of the VAD at 1 Hz, mean flow rate of 3.2 l/min, and ejection pressure of 150 mm Hg. For all tests, the VADs were connected aseptically to a mock loop consisting of a PVC venous reservoir, as previously described.43 The status of the monolayer lining the blood sacs was evaluated by the Alamar Blue assay before and after each run. Our findings for FN and CRY-based coatings are presented in Table 5.2. It is apparent that 24 h of flexing did not dislodge cells from the scaffolds. In fact, the rhythmic movement of the wall even increased cell density markedly in all cultures by more than 13%. This observation is in line with the findings previously reported by others.74 On the other hand, exposure of ECs grown on FN coat-

Fig. 5.4. (a) A typical electron microscopic (SEM) appearance of uncoated, textured PU (magnification x160). The average pore area, evaluated by computer-aided image analysis was 9370 ± 12 µm2. (b) Efficiency of cell attachment onto: A. uncoated PU, and B. CRY-coated PU after seeding with rotation of “ersatz” ventricles. Cell enumeration was performed by Alamar Blue assay and the resulting data were normalized to values obtained for TC. (c,d) Typical SEM views of CRY-coated, textured and smooth PU surfaces, respectively. The specimens were cut at a 30° inclination to reveal both the coating and the underlying PU scaffold (magnification x650).

Tissue Engineering of Prosthetic Vascular Grafts

ing to pulsatile flow resulted in denudation of the monolayers by more than 30%, but did not compromise the monolayers established on CRY coating. In fact, our results indicate that on CRY coating there was minimal, if any, cell loss due to exposure to the hemodynamic forces encountered in a pumping artificial ventricle.

Morphological Aspects of the Endothelialized Surfaces It has long been recognized that the morphology and function of vascular ECs, as well as the architecture of the subendothelial matrix, are modulated by hemodynamic forces and by the tissue-specific microenvironment. In most

Table 5.2. Cell retention under dynamic conditions in EC monolayers established on either FN or CRY coating % of cells remaining % of cells remaining Dynamic Condition on FN coatinga on CRY coatinga Flexing Perfusion Pumping a

120.7 ± 7.6 * 69.0 ± 7.1 * 65.0 ± 16.0 *

113.1 ± 3.3 * 91.6 ± 13.2 95.2 ± 6.3

The data were normalized to the number of cells measured before the test; * indicates significant difference, p 6 mm) positions (e.g., abdominal aortic replacements); however, as the diameter of the synthetic grafts needed became smaller, long term function was compromised. The best evidence for the need of a lumenal endothelial cell lining is provided by the poor patency observed when synthetic grafts with diameters less then 6 mm are used. These grafts exhibit extremely poor, clinically unacceptable, long term patency, due predominantly to the formation of blood clots, resulting in lost patency. The bypass of occluded native vessels with diameters less then 6 mm has been limited to the use of autologous native blood vessels. While these vessels have provided acceptable long term patency, clinicians have realized the importance of maintaining the integrity of the lumenal lining of endothelial cells.1-4 Autologous vessels provide superior long term patency as compared to synthetic grafts, due not only to their compliant natural tissue characteristics, but also to the antithrombogenic nature of the endothelial cell lining. Procedures which disrupt the endothelium, such as angioplasty and atherectomy, provide additional evidence for the need to maintain a functional endothelium. The first report of the successful isolation and transplantation of endothelial cells was a report by Dr. Malcolm Herring in 1978.5 His methods involved the use of a large vesselderived endothelium obtained by scraping the lumenal surface of vein segments. These cells were subsequently transplanted onto the lumenal surface of polymeric graft materials, a process which was termed seeding. The methods used for seeding include mixing the isolated cells with autologous blood and subsequently using this blood-cell inoculum to preclot porous grafts. The lumenal blood-contacting surface was observed to be predominantly fibrin and red cells, with endothelial cells present predominantly within the resulting clot. The hypothesis behind these studies was that seeded endothelial cells would migrate from within the clot to the lumenal blood flow surface, proliferate and finally form a continuous monolayer through the formation of typical interendothelial cell junctions. Subsequent work by a large number of investigators, including work from the laboratories of Schmidt,6 Stanley7 and Graham,8 established the ability to accelerate the formation of monolayers on prosthetic grafts using large vessel endothelial cells as a source of cells for seeding. Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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Publication of these studies established the feasibility of performing endothelial cell transplantation and resulted in extensive work to optimize seeding technology. Numerous questions were identified which needed to be addressed before endothelial cell transplantation was expected to be found clinically acceptable. These questions identified numerous issues related to the source of tissue for subsequent endothelial cell transplantation, the function of immediately isolated endothelium and methods to improve the interaction of endothelium with prosthetic graft surfaces. Subsequent laboratory work has established the efficacy of intraoperative methods for endothelial cell isolation and deposition onto the lumenal surface of vascular grafts. Sources of Endothelium for Transplantation As described above, Herring first reported the use of venous-derived endothelium for subsequent transplantation. The yield of cells from a suitable piece of vein segment was approximately 1 x 104 cells.9 Thus, based on the largest piece of vein which could be obtained from a patient, only 10,000 endothelial cells could be obtained. Under even the most ideal conditions, this cell number would provide a relatively sparse coating of cells, requiring extensive cell proliferation to achieve a monolayer. Kempzinski et al10 reported that even if cells could be isolated and deposited on the surface of polymeric grafts with optimal efficiency, a large number of these cells would be lost from the lumenal surface of the graft due to the effects of blood flow-induced shear on the lumenal surface. Endothelial cells seeded onto grafts with this technology would be required to undergo even more extensive proliferation or, alternatively, a new source of endothelial cells obtained in large numbers must be found.

Cultured Endothelium for Transplantation Methods for the isolation and culture of nonhuman endothelial cells were first reported in the 1920s, providing methods for the expansion of endothelial cell numbers by in vitro proliferation.11 However, the same methods used for nonhuman endothelial cell proliferation had no stimulatory effect on human endothelial cell proliferation. During the 1970s a method for human endothelial cell isolation, based on the pioneering work of Lewis in 1922, provided human endothelial cells in culture; however, these cells exhibited minimal proliferative capacity. A number of endothelial cell growth factors were discovered that slightly stimulated human endothelial cell proliferation. The major breakthrough in human endothelial cell growth in culture was work performed by Dr. Susan Thornton, working with Dr. Elliot Levine, which identified heparin as an integral cofactor with endothelial cell growth factor for the stimulation of human umbilical vein endothelial cell growth in culture.12 Jarrell and coworkers13 subsequently reported that heparin and ECGF would stimulate human adult endothelial cell growth in culture. For the first time, cells from human artery and vein segments from numerous anatomical positions could be grown in large quantities. Large numbers of human endothelia could be produced from single endothelial cells, providing nearly unlimited supplies of human adult endothelium. These breakthroughs in culture

Tissue Engineering of Prosthetic Vascular Grafts

methods provided a means to produce large quantities of autologous endothelial cells by obtaining a small segment of a patient’s blood vessel, isolating and culturing the endothelium in heparin and endothelial cell growth supplement (ECGS) and transplanting these cells at very high seeding densities onto prosthetic grafts. Large vessel-derived endothelial cell seeding at high density was achieved in vitro; however, additional questions have been raised, delaying wide spread clinical utilization of this technique. Concerns related to the use of cultured endothelium for cell transplantations are surmountable. Dr. Peter Zilla has performed a careful series of studies to address the major concerns raised about using cultured cells, as well as providing evidence of the efficacy of this form of tissue engineering.14-16 Some of these questions included concerns relating to the derivation of cells from large vessel sources, which requires a separate surgical procedure, most likely from patients with an already compromised circulatory system. This increased risk of a separate vascular complication had to be weighed against the need for endothelial cell transplantation. If cultured cells were used for transplantation, the conditions of culture must be evaluated to determine the effect on subsequent cell function. Bovine derived serum components such as viruses, and growth factors such as heparin and ECGS, would be carried with transplanted cells into patients, requiring characterization of their effects. Considering that endothelial cells have an extremely low mitotic index, the stimulation of endothelial cell growth must be considered to determine how cell growth may alter endothelial cell function. Finally, little data is available concerning the appropriateness of transplanting vein-derived endothelial cells into an arterial circulation with respect to the ability of these cells to differentiate and function under arterial conditions. Nonetheless, this technique has been successfully implemented in a series of patients undergoing peripheral bypass surgery. Initial results indicate both the safety and efficacy of large vessel endothelial cell transplantation.16

Current Considerations in Endothelial Cell Transplantation Use of Microvessel Endothelium for Tissue Engineering of Vascular Grafts Given the small numbers of cells available if autologous vein-derived endothelium are used without culture, and concerns about culturing effects, investigators have been considering alternate sources of endothelial cells for transplantation. In 1986 Jarrell and Williams17 reported methods for the isolation of autologous microvessel endothelial cells from adipose tissue for use in cell transplantation. The source of fat for this EC isolation was initially fat deposits associated with omentum. This fat is well vascularized, with a density of endothelial cells in excess of 106 endothelial cells per gram of fat isolated. The need to place cells in culture to increase cell number is obviated by the large amounts of endothelium available per gram of omental associated fat. However, access to omental-associated fat still would

Microvascular Endothelial Cell Transplantation: A Review

necessitate a separate surgical procedure, with related complications. The use of omental-associated fat-derived endothelial cells was questioned by Visser et al,18 who reported that omental tissue derived from humans contains predominantly mesothelial cells and not endothelium. This report raised a significant controversy concerning the use of omental-associated fat as a source of endothelial cells for transplantation. Since this initial report by Visser et al, subsequent reports have clarified the controversy and provided additional options for obtaining microvessel endothelial cells.19 First it was established that investigators have erroneously used the terms omentum and omental fat interchangeably. These two tissues are histologically and anatomically distinct. While omentum as used by Visser et al is a vascularized tissue with an extensive number of mesothelial cells present, omental-associated fat, as used by Jarrell and Williams, is composed of predominantly endothelial cells and adipocytes. This is not to say that use of omental-associated fat avoids the possibility of mesothelial cells in the primary cell isolate. Omental fat deposits are covered by a thin layer of mesothelium, and a single cell population isolated from omental-associated fat will contain a small number of mesothelial cells in addition to endothelium. Careful isolation of omental-associated fat away from omentum will reduce contamination by mesothelium. Alternate anatomic sources for microvascularized autologous human fat for endothelial cell transplantation have been reported since these original reports using omentumassociated fat. A significant improvement was the identification of subcutaneous fat as a source of endothelial cells and the use of liposuction to derive this fat from patients.20 A patient undergoing endothelial cell transplantation therefore does not need to undergo a laparotomy to remove omental fat or undergo a separate vascular procedure to remove a segment of vein. A small amount (50 cc) of fat can be removed using a hand held syringe cannula device. This procedure requires less then five minutes and requires a small

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skin incision or trocar puncture to insert the liposuction cannula. The liposuction cannulas used for this procedure have been continuously improved for plastic and cosmetic surgical applications. The use of liposuction fat is also advantageous since the tissue is effectively minced during the liposuction removal process. Scanning electron microscopic evaluation of liposuction fat (Fig. 6.1) illustrates the morphological characteristics of this tissue. The predominant cell type present at this magnification is adipocytes. Recently, a complete characterization of cells present in liposuction-derived fat was reported and established that the major cell type, other than adipocytes, present in this tissue is endothelium.19 Prior to any cell dissociation procedures, human subcutaneous fat contains in excess of 85% endothelium based on the total cells present per unit volume of fat. Adipocytes account for approximately 12% of the cell population, again before any tissue dissociation and cell isolation techniques are used. Therefore, following tissue dissociation and cell isolation, endothelial cells must account for at least 85% of the cells in the final inoculum.

Markers for Isolated Endothelial Cells One significant difficulty in establishing the identity of cells present in the primary isolate following tissue digestion has been the relative lack of a dependable marker of endothelial cells. While antibodies directed against von Willebrand factor or factor VIII-related antigen (FVIIIrAg) have been used somewhat routinely for endothelial cell identification, these markers do not unfailingly react with freshly isolated microvascular endothelial cells. First, upon primary isolation and due in part to the action of proteolytic enzymes, endothelial cells will release these proteins from cytoplasmic granules, thus resulting in cells which exhibit negative staining properties.21 Also confounding is the relative lack of Weibel-palade bodies in fat-derived microvessel endothelial cells. The cytoplasmic inclusions contain extremely high concentrations of vWF, which subsequently stain

Fig. 6.1. Scanning electron micrograph of human liposuction-derived subcutaneous fat, illustrating the predominance of adipocytes.

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extremely well in cell types such as vein-derived endothelium. Venous endothelial cells contain a significant number of Weibel-palade bodies, such that investigators who observe vWF staining in microvessel endothelial cells correctly assess cells with limited reaction product as compared to venous endothelium. Often investigators will suggest that since the bright punctate staining common with Weibel-palade bodies is not observed in microvascular endothelial cells, these cells must not be endothelium. On the contrary, microvascular endothelium is phenotypically distinct from venous or arterial endothelial cells and shows appropriate differences in markers such as vWF and FVIIIrAg. Upon primary culture of microvessel endothelial cells, the morphology often takes on a fibroblastic appearance and, in concert with the relatively minimal reaction with vWF antibodies, investigators will often conclude that the predominant cell type in fat is fibroblasts. What is actually being observed is the ability of endothelial cells to phenotypically differentiate under different in vitro and or in vivo conditions. Interestingly, a recent report suggests that fibroblasts can differentiate in vivo into endothelial cells. This phenotypic drift may be a common characteristic of pluripotent mesenchymal cells, with possible utilization in future cell transplantation technologies. The methods for the isolation of endothelial cells from fat are essentially modifications of the methods first reported by Wagner in 1972 for the isolation of endothelial cells from rat epididymal fat pads.22 Human fat microvessel endothelial cells are isolated by first digesting the fat with a proteolytic enzyme mixture composed predominantly of collagenase and trypsin. The characteristics of this enzyme are critical for successful isolation of endothelial cells from fat.23 Following collagenase digestion the slurry is centrifuged, resulting in the separation of buoyant adipocytes from more dense endothelium. The endothelial cell rich pellet is then used for cell transplantation procedures.

Fig. 6.2. Endothelial cell monolayer established on an ePTFE graft using autologous microvascular endothelial cell transplantation. The graft was implanted as an interpositional device in the carotid artery in a canine model and explanted after 5 weeks.

Tissue Engineering of Prosthetic Vascular Grafts

Cell Deposition on Graft Surfaces: Seeding vs. Sodding The availability of large quantities of autologous endothelial cells does not necessarily overcome another major concern related to endothelial cell transplantation. The deposition of endothelial cells onto the lumenal (blood flow) surface of vascular grafts is a critical step in the transplantation process. Once placed as an interpositional graft, cells on the lumenal surface must resist the flow of blood and remain adherent to the polymeric surface during the maturation of an endothelial cell monolayer. The original methods used to transplant endothelial cells onto graft surfaces were termed seeding procedures, since relatively low concentrations of cells were placed in suspension in either whole blood or plasma, and this suspension transferred to the inner lining of vascular grafts. The endothelial cells in this inoculum were subsequently expected to proliferate and migrate to the blood flow surface, creating a contiguous lining of cells. An alternate form of transplantation, described using the term sodding, was subsequently suggested and evaluated in vitro as well as in preclinical animal and human studies.24 Sodding describes a method of endothelial cell deposition wherein a concentration of cells which would provide a confluent density of cells on the lumenal graft surface is forcibly deposited onto the graft lumenal surface. This forced deposition is most easily carried out by establishing a pressure gradient across the flow surface of the graft whereby the cells are essentially filtered onto the surface of the graft. While previous methods used gravity to allow cell deposition, sodding of cells using pressure deposition results in near immediate association of cells with the lumenal surface. The time to prepare a tissue engineered vascular graft for implantation, a period inclusive of tissue isolation, cell isolation and cell deposition, has effectively been reduced to less than 60 minutes.20

Microvascular Endothelial Cell Transplantation: A Review

Animal Models of Endothelial Cell Transplantation

97

Mechanisms Underlying the Formation of a Neointima in Tissue Engineered Vascular Grafts Using Microvascular Endothelial Cell Transplantation

formation of an endothelial cell lining, but questions have been raised regarding whether the cells present on the lumenal surface of microvessel sodded grafts represent the same cells transplanted during the initial sodding procedure. This question has been addressed using a method to trace the fate of transplanted endothelial cells using a fluorescent tagging method.35 A fluorescent dye is permanently incorporated into the membranes of microvessel endothelial cells at the time they are transplanted onto the lumenal surface of vascular grafts. This dye has unique characteristics in that it is long lived and provides a marker of the fate of transplanted cells. If, however, fluorescently labeled cells duplicate, each daughter cell receives one half the fluorescence of the initial cell. For cells which undergo multiple duplication, a process which could occur following cell transplantation, the subsequent progeny would contain undetectable amounts of fluorescence. Use of this technique in a canine model of microvessel endothelial cell sodding revealed that after 5 weeks the monolayer of endothelial cells on the flow surface exhibited extensive fluorescence, establishing their identity as cells present in the original cell inoculum used for sodding. Thus, following sodding, the microvessel endothelial cells need simply adhere to the polymer surface and to each other to form a confluent monolayer without the need to undergo extensive cell proliferation.

A number of preclinical animal studies have established that, following the transplantation of microvessel endothelial cells onto synthetic grafts, the accelerated formation of an endothelial cell monolayer on the lumenal flow surface is observed. The earliest time point evaluated is 4 days in a rat aortic graft model of microvessel endothelial cell transplantation.34 At this time a confluent layer of endothelial cells is observed. Studies in dogs have established that the endothelial cell lining observed in microvessel endothelial cell sodded grafts remains stable for periods of up to one year, the longest published animal implant data.33 Figure 6.3 illustrates the lining on a microvascular endothelial cell sodded ePTFE graft explanted after one year. This morphologic data establishes the ability to accelerate the

Preclinical Animal Trials of Endothelial Cell Transplantation With the development of methods for the isolation of endothelial cells from numerous tissue sources, including macro- and microvascular sources, numerous animal studies have been performed to evaluate the effects of endothelial cell transplantation on small diameter vascular graft patency. These studies use small diameter grafts (< 6 mm), most often in the canine species. The predominance of canine models for these preclinical evaluations is based on the general acceptance of this model as being similar to humans with respect to graft healing. While the canine species is commonly used, techniques for graft placement (e.g., end to end,

The availability of methods for endothelial cell isolation and culture, as well as methods to deposit cells onto graft surfaces, has resulted in numerous investigations to evaluate the efficacy of endothelial cell transplantation.25-33 The earliest animal models were developed to evaluate the ability to accelerate the formation of endothelial cell linings on polymeric grafts. Most of these studies have been performed in the canine model, due primarily to the acceptance of this animal model as a predictor of graft function in humans. Earliest studies established the ability to accelerate the formation of a continuous monolayer of endothelial cells on the lumenal surface of vascular grafts (Fig. 6.2). In general, monolayer formation is highly dependent on the density of cells used to treat the graft surface. When cells are transplanted at confluent densities or greater, monolayer formation appears to be complete in under three weeks. Subconfluent densities require more extended periods of time.

Fig. 6.3. Scanning electron micrograph of a sodded ePTFE carotid artery after one year.

*For chi-square >3.84 for 1 degree of freedom, P τ) to this strain increment (held constant after its application) is

σ(t ) ∆ε (τ) = G (t − τ )

dε τ ∆t . dt

(21)

Now, the principle of linear superposition allows us to sum the elastic responses to all the past increments in load from time -∞ to t. By taking infinitesimal increments in time we can write Equation (21) in integral form as

σ( t) =

∫

t

dε dτ . dt

G ( t − τ)

−∞

(22)

Integrating by part yields

σ( t) = G(0 ) ε (t ) −

t

dG(t − τ) ε (τ )dτ dt -∞

∫

(23)

where G(0) is the so called glassy or instantaneous modulus. Similarly, the delayed strain response to a smooth stress function of time is

ε (t) = J ( 0) σ(t ) − where

d J (t − τ) σ( τ) dτ dτ −∞

∫

J (t ) =

t

∫

(24)

∞

T(τ)(1 − e −t τ )dτ ,

and

0

T(τ) is the retardation spectrum. Equations (23) and (24) are constitutive laws for a linear viscoelastic material. They can be implemented in a numerical algorithm, such as the finite element method, to predict the history of any geometry with a continuum viscoelastic material behavior subjected to a time-varying stress or strain. Modeling Microstructure It was noted earlier that the complete geometry of the structure must be fully characterized. However, in the case of a porous material, it would neither be possible, nor desirable, to model precisely the structure of each pore. There-

449

fore, in this case it would be advantageous to include information about the microstructure in the description of the material behavior. The porous microstructure is then modeled as a continuum (solid) region that has the apparent (or effective) properties of the porous material. This idea can be taken one step further by including structural properties, such as porosity, as variables in the material formulation. Obviously the relation between the material behavior and the structural property needs to be determined. But once this has been achieved, it becomes a simple matter to examine the influence of changes in material microstructure on the structural response of the component without having to undertake complex geometrical modeling. In most porous materials the single most important parameter in determining the mechanical properties is the relative density ρ/ρs, where ρ is the apparent density while ρs is the density of the structural members. The porosity is therefore defined as 1-ρ/ρs. It turns out that cell size is generally far less important in the determination of material properties; however, cell shape will have a notable influence. The distinction needs to be made between open and closed cells: In an open configuration, the cells are interconnected, whereas in a closed cell arrangement each cell is isolated from its neighbor. As interconnectivity is vital in a tissue-engineered graft, only the mechanics of open celled structures will be considered. Most porous structures are anisotropic to a greater or lesser extent. Anisotropy arises from the shape of the cells, and is characterized by the ratio of the largest to the smallest cell dimension (R). Typically in a foam that is considered to be isotropic, R ≈ 1.3; however, values in the range from one to ten are not uncommon. For most linear-elastic, isotropic, cellular structures the apparent material modulus E can be related to the apparent density by

E E s ∝ (ρ ρ s )

M

(25)

where Es is the elastic modulus of the solid material, and M is a constant which is generally greater than one. However, if the material structure is anisotropic, and therefore the stiffness cannot be represented by a scalar quantity, a more complex description is required. Common porous materials have distinct directionality and can be regarded as either transversely isotropic or orthotropic. A transversely isotropic material has one preferred direction and any transverse plane will be isotropic. An orthotropic material will have three orthogonal planes of symmetry. For these materials the moduli in the preferred directions will be a function of the density as well as the anisotropy ratio. For most porous materials the elastic modulus is approximately proportional to R2 while the strength is typically proportional to R.47 Fig. 40.7. Standard linear viscoelastic models.

450

Based on analytical and experimental studies,47 parameters such as the yield strength and the compressive collapse load of a porous material can also be determined as a function of relative density and material structure. However, these investigations are limited to the case where the material comprising the structure of a foam is operating in the linear elastic regime. In a highly porous material subjected to large forces this would not be the case, and therefore other means of characterizing the material response are required. One approach for determining the apparent properties of a material is to develop numerical models of the microstructure under investigation. If the behavior of the material from which the structure is made has been completely characterized, and if the material has a uniform structure, then a single structural unit can be tested in order to obtain the apparent properties of the material. An example of this approach is shown in Figure 40.8. The polyurethane microstructure (Fig. 40.8a) has been idealized and modeled using the finite element method (Fig. 40.8b). This model is then subjected to a range of numerical tests, from which the apparent material behavior can be determined. This approach has great flexibility, as it makes possible the investigation of the influence of relative density, cell shape and size. It also becomes a relatively simple matter to distort the cells so as to create an anisotropic structure and then quantify the structural response. Once the nature of the apparent material has been determined this information can be used to model a complete implant (Fig. 40.8c) without having to consider the geometry of the internal microstructure. The presence of a viscous fluid within a porous material will have a considerable influence on the strength. When the material is compressed the fluid will be squeezed out, while it will be drawn in when the material is extended. As the fluid has viscosity, work is done during material deformation. The higher the loading rate the greater will be the

Fig. 40.8. (a) Micrograph of a polyurethane microstructure and (b) and idealized model of the microstructure which is then used (c) for the determination of an equivalent material for the properties of a porous structure.

Tissue Engineering of Prosthetic Vascular Grafts

amount of work, and therefore a strong dependence on strain rate exists. In the early postoperative stages of a tissue-engineered graft, the graft material will contain fluid. In addition, the normal loading rate of approximately one cycle per second will likely be high enough to make the influence of loading rate noticeable. For these reasons the fluid phase of the porous structure needs to be given consideration. The foam can be treated as a porous material with an absolute permeability K. The velocity of the fluid v within the porous material can then be determined from Darcy’s law

v= −

K dp , µ dx

(26)

where µ is the dynamic viscosity of the fluid and dp/dx is the pressure gradient within the porous structure. The permeability can be determined from the relative density and the pore structure, while the pressure gradient can be related to the applied stress. After some mathematical manipulation the strength of the porous material can be related to the viscosity of the fluid, the strain rate and some geometrical information. Material Testing Material testing forms the basis of the characterization of the mechanical properties of the native artery and the graft material. The material morphology and the extent to which the material needs to be characterized determine the complexity of the testing. For isotropic materials, simple uniaxial extension tests may be sufficient to obtain the necessary information for characterizing the material. However, for transversely isotropic and orthotropic materials, biaxial testing is required. Hayashi et al9 developed a highly elastic, blood compatible, small diameter vascular graft fabricated from segmented polyurethane. Introducing porosity into the material, which also allowed for possible tissue ingrowth, created

An Integral Mathematical Approach to Tissue Engineering of Vascular Grafts

a physiologically realistic compliance for the graft. Electron microscopy revealed the graft material to be approximately isotropic, and therefore the mechanical properties were determined using simple uniaxial tests. Their strength tests revealed the importance of pretreatment of the material prior to mechanical testing. They found that the material became stable, in terms of mechanical properties, after having been immersed in a saline solution for 30 minutes. They compared the stress/strain characteristics of their graft material to those of traditional grafts as well as the natural artery (Fig. 40.9). In order to present excessive dilation, the graft was coated with a woolly polyester net. However, such a net will reduce the compliance of the graft. The mechanical test results showed that the porous polyurethane exhibited an almost linear stress/strain response (KP), with a slight tendency to strain more at higher stresses. This result is in contrast to the highly non-linear behavior of the calf ’s thoracic aorta (Ao) which shows a very strong decline in the amount of strain at elevated stresses. However, in the low strain region (up to approximately 30%) the stiffness of the polyurethane graft is very similar to the thoracic aorta. This result further shows the considerable difference in stiffness between the compliant polyurethane graft and the more stiff traditional Dacron grafts (CW, GT, CV and DB). Although the material developed for this graft is very compliant and exhibits strain of over 200% before failure, it does not show the stiffening behavior normally associated with polymeric materials. The slight softening at high strains may be a result of the failure of the woolly coating. Further-

451

more, the time dependent behavior (viscoelasticity) of the material will influence the sizing of the graft, and therefore needs to be considered in the design of the implant. Hyperelastic Properties In order to develop the constitutive models for soft tissues and compliant engineering materials, the general form of the strain energy function must be postulated. This should be as simple and convenient as possible, as laboratory tests will be required to set the relevant parameters in the strain energy function. These energy functions are based on qualitative observations of the material behavior and on prior testing of similar materials. The functional form of the strain energy function is subject to much debate and both polynomial and exponential forms are in current use. The development of sound constitutive laws to describe the mechanical behavior of arterial vessels has been the focus of numerous research efforts.17,48-50 Since these constitutive laws are set out in terms of large strain hyperelasticity, experiments are required to determine both the functional form and the coefficients defining the strain energy function, W. Essentially three forms of strain energy functions are used to describe arterial walls: polynomial, exponential and logarithmic. The polynomial function of Vaishnav et al,51 which was used to describe a cylindrical vessel subject to internal pressure, is the simplest; in general it has seven coefficients and is given as follows: (27) Fig. 40.9. Comparison of stress— strain characteristics for a porous polyurethane graft (KP) with other traditional graft materials (CW—Cooley low porosity woven Dacron, GT— Gore-Tex thin walled EPTFE, CV— Cooley double velour knitted Dacron) as well as the thoracic aorta (Ao). Adapted from Hayashi et al 1989.

452

Tissue Engineering of Prosthetic Vascular Grafts

where a, b, c, d, e, f, g are material constants, and Eθθ and Ezz are the longitudinal strain components in the circumferential and axial directions, respectively. Fung et al17 proposed an exponential function of the form

(

2 W = C exp a1 Eθθ + a 2 E 2zz + a4 Eθθ E zz

)

(28)

where c, α1, α2 and α4 are the material constants. Experimental data was determined from the cardiac, left iliac, lower aorta and upper aortic arteries of rabbits, which was used to determine the coefficients defining the strain energy functions. The comparison of the experimental and theoretical results, using Equation (10), are shown in Figure 40.2. Lastly, Takamizawa and Hayashi18 proposed a logarithmic function that again has four coefficients

[

]

2 W = − C ln 1 − 12 a θθ Eθθ − 21 a zz E zz2 − a θ z Eθθ E zz (29)

where C, aqq, azz and aqz are material constants. The different layers and components making up the arterial wall result in an inhomogenous structure—that is to say, the properties of the material change through the thickness of the wall.52 The models that have been previously discussed have not included the different layers when analyzing the mechanical response. To address this limitation, von Maltzahn et al53 modeled the artery using two layers, each having different properties. The arterial wall is generally considered to be anisotropic; however, this does not necessarily mean that each layer within the wall is anisotropic. Their model consisted of a transversely isotropic media and an anisotropic adventitia. This two layer model was compared to the single layer model of Vaishnav et al,51 and it was found that it more accurately predicted the experimental data. This was the case even when the fourth order terms were included into the polynomial function. The model also had the advantage that the material parameters could be related to specific mechanical properties within the artery. Viscoelastic Properties The appropriate choice of experiment for the determination of viscoelastic properties depends on a number of factors. If the modulus or compliance function is to be approximated with only a few relaxation times, then either a stress relaxation or a retarded strain test would suffice. The time constants and associated moduli or compliance coefficients would be found by curve fitting to the assumed form of the rheological model. To characterize the full constitutive behavior, tensile, shear and volumetric (if the material is compressible) testing must be done. Such time-history tests are problematic in practice because they require the application of sudden large stresses or strains without inducing significant inertial effects. This means that the loading cannot be applied too quickly and so very small relaxation time scales cannot be measured with such tests. The determination of very large time constants—longer than a few days— is also impractical. The long term behavior of materials which do creep (such as non-crosslinked polymers) can thus not realistically be fully characterized.

Dynamic testing with sinusoidal loading over a large range of frequencies will provide data for the determination of the continuous spectrum functions S(τ) and T(τ). The viscoelastic strain response to a sinusoidal stress,

σ ( t ) = σ 0 sin ω t ,

(30)

at frequency ω, and amplitude σ0 is

ε( t ) = ε 0 sin(ω t - δ )

(31)

where the peak strain ε0 lags the peak stress by δ—the loss angle. These relations can be written in complex form as

σ ( t) = σ *0 e iω t and ε(t ) = ε*0e

i(ω t −δ )

(32)

where ρ*o and ε*0 are the complex amplitudes. The complex modulus, made up of a real (stiffness) and complex (damping) part, is σ 0 iδ e = G1 + iG 2 (33) ε0 with an amplitude ε0 lagging the stress. The loss tangent is the ratio of the complex to real parts of G*, viz.

G∗ =

tanδ =

G2 . G1

(34)

Thus, measuring ρo, ε0, and δ for a range of frequencies and using Equations (33) and (34), we can plot G1 and G2 as functions of frequency. An appropriate continuous relaxation spectrum S(τ) is chosen based on the form of these curves. Fung22 describes this procedure in detail. See also Ferry54 for a comprehensive treatise on this topic. Fung24 found that for a wide range of frequencies, the damping modulus is constant for biological soft tissues. This result allows for a very simple relaxation spectrum, S(τ)=c/τ. Lockett 55 reported the same result for polymethylmethacrylate (PMMA). However, he found a considerable frequency dependency above 1 Hz for a lightly crosslinked amorphous polymer. One other important consideration when undertaking materials testing is the fact that viscoelastic properties tend to be highly temperature dependent. There are established, proven laws which allow one to simply shift the compliance/ time curve for a particular material up or down the time line if the temperature is decreased or increased respectively. In the case of tissue engineered products this is not an issue for developing a constitutive model. However, it is vital that all materials tests are done at normal core body temperature. A detailed treatment on the issues involved in viscoelastic materials testing is given by Lockett.55 Other useful literature includes Christensen56 and Sternstein.57

Finite Element Implementation General Outline The finite element method is a numerical procedure frequently used for analyzing complex deformable bodies

An Integral Mathematical Approach to Tissue Engineering of Vascular Grafts

subjected to mechanical loading. The finite element procedure essentially involves discretizing a component into a number of small (finite size) regions called the elements, (which are defined by nodes—or nodal points), and approximating the mechanical response in each element by assuming that the local deformation can be represented by loworder polynomials. For time independent linear problems, the potential energy of the entire system can be minimized in one step that produces a large number of algebraic simultaneous equations in the nodal displacements, which are solved using a computer. The results obtained using this method are not exact but with increased model refinement, and subsequently more computational effort, solutions to the approximated displacement field of the desired resolution may be obtained. For the non-linear and time-dependent problems presented in the design of artery grafts, the problem is broken down into a number of linearized increments. The method therefore requires repeated solution of simultaneous equations to solve for increments in nodal displacements. However, to illustrate the finite element method here, we limit the discussion to the linear, time independent problem. The finite element method requires: 1. That the behavior of the component material be precisely defined; 2. That the geometry of the component in the unloaded state be known; and 3. That the boundary conditions, which include the mechanical loading, the supports or restraints holding the structure and, for time dependent problems, the initial velocities, are fully characterized. From this problem definition the finite element procedure will be able to analyze the structure and calculate the design parameters, such as the deformations and stresses over the entire component domain. These results will allow decisions on the appropriate material choice and component geometry to be made. The flexibility and predictive power of the finite element method allows engineers the ability to test various design parameters. In the case of a porous graft, for example, the finite element method is an efficient way to examine the effect on overall component compliance of changes in the graft diameter, thickness, material stiffness, porosity, pre-stress, etc. Minimizing the Potential Energy In the continuum description of hyperelastic material behavior, it was most convenient to use tensor notation to describe the field quantities (stresses, strains and displacements) in a deformable body. However, since the finite element method describes these fields with their values at discrete points (for example, at the nodes for the displacements) it is most effective to use matrix notation for the finite element equations. Since the finite element method computes the displacement field that minimizes the potential energy, the primary variables are the displacements of the nodes. The vec-

453

tor {u}represents these displacements. The finite element method seeks to minimize the potential energy by selecting the values of {u} which minimize

U({u}) = W (E ({u}))dV − {f } {u} T

∫

(35)

V

for all kinematically admissible {u}, where V is the volume of the body and {f} is the vector of nodal loads. Element Behavior To calculate elastic strain energy in the body, the material domain is discretized into small regions called elements (see, for example, Fig. 40.10). The geometry of each element is described by the location of a set of nodes (different element types have different numbers of nodes) which also serve as inter-element connections. Within each element the displacement field is approximated by low-order piecewise continuous polynomials, specially chosen such that their coefficients are the nodal displacement components for that element, {αe}, which is a subset of the global displacements, {u}. In each element, the element strains, {E}, are obtained from the displacements of the nodes in that element {αe}, by writing Equation (3) as the compatibility requirement

[

]

{E} = B({a e }) {a e }.

(36)

The Piola-Kirchhoff stress components in each element are written as a vector {S}, so that the constitutive relation, Equation (10), which accounts for all of the interactions between the stress and strain components in a linear elastic body, can now be written as

{S }= [D]{E}

(37)

where [D] is the square matrix of coefficients representing the tensor Dijkl in Equation (12). Essentially, the matrix [D] describes the stresses required to produce the given strains in the material, and is determined by the strain energy function, W, using Equation (12). The strain energy W is determined by laboratory testing on the material of interest. Materials such as soft tissue exhibit non-linear time dependent behavior and Equation (13) must be written in an incremental form

{dS }=

[D({E}, {E& })]{dE}

.

(38)

Therefore the components of [D] depend on the strain {E}and the strain rate {E}. Materials of this type can cause significant computational difficulties, but the general procedures are the same as in the linear case. For time-independent hyperelastic materials, the elastic strain energy in each element, We, is computed by integrating the hyperelastic strain energy potential W over the volume of the element, Ve, viz.,

454

Tissue Engineering of Prosthetic Vascular Grafts

∫ = ∫

We =

displacement field, can be used to compute the nodal displacements, {αe}, from which the strains, {E}, and the stresses, {S}, can be computed within each element. Hence the distribution of stress and strain can be found over the region of interest.

W dV e

Ve

1 2

=

1 2

T

Ve

{S } {E}dV e T

{a}

∫

V

T

e

[B] [D][B]dV e {a}

(39)

where we have used both the material constitutive equation (37) and the compatibility relationship (36) for that element (note that the superscript ( )T denotes the transpose of the vector or matrix). It is convenient to define the element stiffness matrix [ke], by T

[k ]= ∫ [B] [D][B]dV e

e

(40)

Ve

so that the elastic potential energy (or strain energy) for a single element is given by T

({ })= {a } [k ]{a }

W e ae

1 2

e

e

e

(41)

Element Assembly To compute the strain-energy for the entire structure, WΣ({u}), the element strain energies, We({αe}), are summed. This is done by summing the element stiffness matrices, [ke], for each element. Since each set of element displacements, {αe}, is a different subset of the global displacements, {u}, we must relate the local to the global displacements, which can be done by defining a locator matrix for each element, from which {αe}=[Pe]{u}. This allows us to compute a global stiffness matrix, which represents the structural stiffness of the entire component. Mathematically,

(42) Problem Solution To compute the total energy of the system, U({u}), we require the work done by the loads on the component, which is merely the inner product of the loads at each node, {f}, and the nodal displacements, {u}. Thus, the total energy of the system

U ({u}) =

1 2

{u}T [K ]{u}− {u}T {f }.

(43)

Minimizing the total energy requires the solution of the (usually large) system of simultaneous linear equations

[K ]{u}− {f } = {0},

(44)

for which modern computers are highly effective. The displacements, {u}, which are an approximation to the true

Examples of Artery Graft Modeling Once the mechanics of a material has been fully characterized, the geometry of the graft can be modeled using the finite element method. To complete the modeling process, the loading and boundary conditions need to be described in terms of the finite element modeling. In the case of the natural artery, or a vascular graft, the loading magnitudes are easily quantified from blood pressures. In addition, the time dependence of the loading pulse can also be determined and used as input to the numerical model. In order to illustrate some aspects of the mechanical design process, a few examples of artery graft modeling are presented. Figure 40.10 shows the displaced shape of a simple finite element model of the anastomosis region of the natural artery and a porous polyurethane graft. The artery (shaded) is modeled as a single layer, using the simple hyperelastic material model proposed by Fung et al,17 and does not include any pre-stress. The graft is modeled as a hyperelastic and isotropic material, using the experimental data from the uniaxial tensile test data of the porous polyurethane. The graft is analyzed for a range of relative porosities. The different porosities were created by altering the ratio of the salt to polymer concentration in the manufacturing process (values of 1:1, 3:1, 5:1 and 7:1 have been investigated). Both the native artery and the graft are assumed to have an internal diameter (D) of 6 mm and a wall thickness of 1 mm. The model was loaded with an internal pressure, which is initially set to 70 mm Hg, and then increased linearly to 120 mm Hg—thus ∆P is 50 mm Hg. The resultant compliance of the artery and the graft, as a function of the axial distance, is shown in Figure 40.11. A simple measure of vascular compliance58

 2  ∆ D  Cv =    ,  D  ∆ P 

(45)

which relates the change in graft diameter ∆D to the change in pressure, is used to compare the compliance of the native artery to the porous graft. This relation is chosen, as it does not contain any information about the properties of the two materials, which are non-linear and therefore cannot be characterized with a single constant parameter. The compliance of the native artery was predicted to be 0.273%/mm Hg (or 13.65% for a pulse pressure of 50 mm Hg), which was greater than any of the four grafts analyzed. In the region of the anastomosis the compliance of the artery is affected by the presence of the graft and therefore reduces towards the junction, which is halfway (5 mm) along the length. The lowest porosity graft (1:1) results in the smallest compliance, which is 0.05%/mm Hg. As the porosity increases to 7:1, the graft compliance increases to approximately 80% that of the host. It is interesting to see that with this more compliant graft material a hyper-com-

An Integral Mathematical Approach to Tissue Engineering of Vascular Grafts

pliant zone of about 2 mm, on the arterial side of the anastomosis, is predicted. In this region the compliance of the artery is increased by approximately 0.4%. A similar zone of hyper-compliance was predicted by Chandan et al59 and observed in vivo by Hasson et al.60 However, in the in vivo study the extent of hyper-compliance was more dramatic, with a localized increase in compliance of up to 50%. The graft compliance as a function of the porosity is shown in Figure 40.12. The result shows that the percentage compliance increases exponentially as the porosity increases. However, increase in porosity will also be accompanied by a decrease in yield strength of the material.

455

The results of the proceeding example are based on a static analysis and do not show the influence of the time dependent behavior of the material. The viscoelasticity of the polyurethane will result in a gradual increase in the graft diameter, as a result of strain accumulation during cyclic loading. The increase in diameter will reach a steady state after some time. Therefore the inclusion of viscoelasticity is important in determining the final size of the graft after implantation. The flexibility of the modeling approach means that complex geometrical features, with a host of different interacting materials, can be simulated. Figure 40.13 shows

Fig. 40.10. Simple axisymmetric finite element model of anastomosis region of the native artery and the polyurethane graft.

Fig. 40.11. Compliance (%/mmHg) of the native artery and the polyurethane graft for various porosity’s of the graft material shown as a function of the axial distance.

456

Tissue Engineering of Prosthetic Vascular Grafts

a finite element model of a graft that has been reinforced with two spiral wound fibres. This type of model can be used to determine the compliance and the stresses within the reinforcing fibres, as well as the local strain distribution through the wall of the graft material.

Conclusions The design of a tissue-engineered vascular graft involves mechanical issues that are ideally suited to sophisticated analysis methods such as the finite element method. However, the mechanics of both the intact artery and the graft materials is extremely complex. The complexities arise from the many different characteristics that the materials exhibit, such as the non-linearity, hyperelasticity, viscoelasticity, anisotropy. Combining all these three characteristics, while also introducing the effect of the microstructure, results in an involved material formulation. However, once the materials involved have been adequately characterized, a host of design issues can be examined with relative ease. The finite element method is therefore an ideal tool for aiding the development of the mechanical aspects of a tissue-engineered vascular graft.

References 1. Rodgers VGJ, Teodori MF, Borovetz HS. Experimental determination of mechanical shear stress about an anastomotic junction. J Biomech 1987; 20:795-803. 2. Weston MW, Rhee K, Tarbell JM. Compliance and diameter mismatch affect the wall shear rate distribution near an end-to-end anastomosis. J Biomech 1996; 29:187-198. 3. Stewart SFC, Lyman DJ. Effects of vascular graft/natural artery compliance mismatch on pulsatile flow. J Biomech 1992; 25:297-310. 4. Fry DL. Acute vascular endothelial changes associated with increased blood velocity gradients. Circ Res 1968; 22:165-197. 5. Wong AJ, Pollard TD, Herman IM. Actin filament stress fibres in vascular endothelial cells in vivo. Science 1983; 219:867-869. 6. Ishibashi H, Park H, Ojha M, Langdon S, Langille L. Shear-intimal thickening on the bed of an end-to-end anastomosis model. Advances in Bioengineering ASME BED- 1996; 33:475-476. 7. Lyman DJ, Fazzio FJ, Voorhees H, Robinson G. Compliance as a factor affecting the patency of a copolyurethene vascular graft. J Biomed Mater Res 1978; 12:337-345. 8. Seifert KB, Albo D, Knowlton H, Lyman DJ. Effect of elasticity of prosthetic wall on patency of small-diameter arterial prostheses. Surg Forum, 1979; 30:206-208.

Fig. 40.12. Graft compliance as a function of the material porosity (salt:polymer ratio).

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Fig. 40.13. Finite element model of a possible graft structure showing an underlying porous graft material surrounded by a reinforcing weave.

9. Hayashi K. Elastic properties and strength of a novel small-diameter compliant polyurethane vascular graft. J Biomed Mater Res 1989; 23:229-244. 10. Deweese JA. Anastomotic neointimal fibrous hyperplasia. In: Complications in vascular surgery, 2nd edition. Orlando: Grune and Stratton, 1985:157-170. 11. Ballyk PD, Ojha M, Walsh C, Butany J. Suture-induced intramural stresses and intimal hyperplasia. Advances in Bioengineering ASME BED- 1996; 33:213-214. 12. Ballyk PD, Walsh C, Ojha M. Effect of intimal thickening on the stress distribution of an end-to-side graft-artery junction. Advances in Bioengineering ASME BED- 1995; 31:327-328. 13. Paasche PE, Kinley CE, Dolan FG, Gonza ER, Marble AE. Consideration of suture line stresses in the selection of synthetic grafts for implantation. J Biomech 1973; 6:253-259. 14. Greer LS, Vito RP, Nerem RM. Material property testing of a collagen-smooth muscle cell lattice for the construction of a bioartificial vascular graft. Advances in Bioengineering ASME BED 1994; 28:69-70.

15. Fung YC, Biomechanics: Circulation 2 nd Edition, Springer,1996. 16. Young JT, Vaishnav RN, Patel DJ. Non-linear anisotropic viscoelastic properties of canine arterial segments. J Biomech 1977; 10:549-559. 17. Fung YC, Fronek K, Patitucci P. Pseudoelasticity of arteries and the choice of its mathematical expression. Am J Physiol 1979; 237:H620-H631. 18. Takamizawa K, Hayashi K. Strain energy density function and uniform strain hypothesis for arterial mechanics. J Biomech 1987; 9:293-300. 19. Cox RH. Anisotropic properties of the canine carotid artery in vitro. J Biomech 1975; 8:293-300. 20. Dobrin PB. Biaxial anisotropy of dog carotid artery: Estimation of circumferential elastic modulus. J Biomech 1986; 19:351-358. 21. Tanaka TT, Fung YC. Elastic and inelastic properties of the canine aorta and the variation along the aortic tree. J Biomech 1974; 7:357-370. 22. Fung YC. Biomechanics: Mechanical properties of living tissue 2nd Edition, Springer,1993.

458 23. Chuong CJ, Fung YC. Compressibility and constitutive equations of arterial wall in radial compression experiments. J Biomech 1984; 17:35-50. 24. Fung YC. Stress-strain history relations of soft tissues in simple elongation. In: Fung YC, Perrone N, Anliker M, eds. Biomechanics: Its foundations and objectives Englewood Cliffs: Prentice-Hall, 1972. 25. Hutchinson KJ. Effect of variation of transmural pressure on the frequency response of isolated segments of canine arotid arteries, Cirulation Res 1974; 35:742-751. 26. Newman DL, Bowden NLR, Gosling RG. The dynamic and static elastic properties of the aorta of the dog. Cardiovasc Res 1975; 9:679-684. 27. Greenwald SE, Newman DL, Denyer HT. Effect of smooth muscle activity on the static and dynamic elastic propertries of the rabbit carotid artery. Cardiovasc Res 1982; 16:86-94. 28. Langewouters GJ, Wesseling KH, Goedhard WJA. The static elastic properties of 45 human thoracic and 20 abdominal aortas in vitro and the parameters of a new model. J Biomech 1984; 17:425-435. 29. Langewouters GJ, Wesseling KH, Goedhard WJA. The pressure dependent dynamic elasticity of 35 thoracic and 16 abdominal human aortas in vitro described by a five component model. J Biomech 1985; 18:613-620. 30. Haut RC, Little RW. A constitutive equation for collagen fibres. J Biomech 1972; 5:423-430. 31. Dehoff PH. On the nonlinear viscoelastic behavior of soft biological tissues. J Biomech 1978; 11:35-40. 32. Decraemer WF, Maes MA, Van Huyes VJ, Van Peperstraete P. A non-linear viscoelastic constitutive equation for soft biological tissues. J Biomech 1980; 13:559-564. 33. Wu SG, Lee GC. On nonlinear viscoelastic properties of arterial tissue. ASME J Biomech Eng 1984; 106:42-47. 34. Demiray H. A quasi-linear constitutive relation for arterial wall materials. J Biomech 1996; 29:1011-1014. 35. Johnson GA, Livesay GA, Woo SL-Y, Rajagopal KR. A single integral finite strain viscoelastic model of ligaments and tendons. J Biomech Eng 1996; 118:221-226. 36. Lanir Y. A microstructural model for the rheology of mammalian tendon. J Biomech Eng 1980; 102:332-339. 37. Lanir Y. Constitutive equations for fibrous connective tissue. J Biomech 1983; 16:1-12. 38. Minns RJ, Soden PD, Jackson DS. The role of the fibrous components and ground substance in the mechanical properties of biological tissues: A preliminary investigation. J Biomech 1973; 6:153-165. 39. Huyghe JM, Van Campen DH, Arts T, Heethaar RM. The constitutive behavior of passive heart muscle tissue: A quasi-linear viscoelastic formulation. J Biomech 1991; 24:841-849. 40. Myers BS, McElhaney JH, Doherty BJ. The viscoelastic responses of the human cervical spine in torsion: Experimental limitations of quasi-linear theory, and a method for reducing these effects. J Biomech 1991; 24:811-817. 41. Chuong CJ, Fung YC. On residual stresses in arteries. J Biomech Eng 1986; 108:189-192.

Tissue Engineering of Prosthetic Vascular Grafts 42. Fung YC, Liu SQ. Strain distribution in small blood vessels with zero stress state taken into consideration. Am J Physiol 1992; 264:H544-H552. 43. Greenwald SE, Rachev A, Moore JE, Meister JJ. The contribution of the structural components of the arterial wall to residual strain. Advances in Bioengineering ASME BED 1994; 28:63-64. 44. Humphrey JD, Strumpf RK, Yin FCP. Determination of a constitutive relation for passive myocardium: I. A new functional form. J Biomech Eng 1990; 112:333-339. 45. Humphrey JD, Strumpf RK, Yin FCP. Determination of a constitutive relation for passive myocardium: II.— Parameter estimation. J Biomech Eng 1990; 112:340-345. 46. Van Dijk AM. Vascular single smooth muscle cells and whole tissue. Mechanical properties and response of stimuli. PhD Thesis University of Leiden, Netherlands: 1983. 47. Gibson LJ, Ashby MF. Cellular solids: Structure and properties, Pergamon Press, 1988. 48. Carmines DV, McElhaney JH, Stack R. A piece-wise nonlinear elastic stress expression of human and pig coronary arteries tested in vitro. J Biomech 1991; 24:899-906. 49. Hayashi K. Experimental approaches on measuring the mechanical properties and constitutive laws of arterial walls. J Biomech Eng 1993; 115:481-487. 50. McAfee MA, Kaufmann MV, Simon BR, Baldwin AL. Experimental/numerical approach to the determination of material properties of large arteries. Advances in Bioengineering ASME BED 1994; 28:111-112. 51. Vaishnav RN, Young JT, Patel DJ. Distribution of stresses and of strain-energy-density through the wall thickness in a canine aortic segment. Circ Res 1973; 32:577-583. 52. von Maltzahn WW, Warriyar RG, Kietzer WF. Experimental determination of the elastic properties of media and adventitia of bovine carotid arteries. J Biomech 1984; 17:839-847. 53. von Maltzahn WW, Besdo D, Wiemer W. Elastic properties of arteries: A non-linear two-layer cylindrical model. J Biomech 1981; 14:389-397. 54. Ferry JD. Viscoelastic properties of polymers. New York: Wiley and Sons, 1961. 55. Lockett FJ. Nonlinear Viscoelastic Solids. New York: Academic Press, 1972. 56. Christensen RM. Theory of Viscoelasticity: An Introduction, New York: Academic Press, 1971. 57. Sternstein SS. Mechanical properties of glassy polymers: Treatise on Materials Science and Technology. In: Schultz JM, ed. Volume 10, part B, New York: Academic Press 58. Gow BS, Taylor MG. Measurement of viscoelastic properties of arteries in the living dog. Circ Res 1968; 23:111-122. 59. Chanran KB, Gao D, Han G, Baraniewski H, Corson JD. Finite element analysis of arterial anastomoses with vein, Dacron and PTFE grafts. Med Biol Eng Comp 1992; 30:413-418. 60. Hasson JE, Megerman J, Abbott WA. Increased compliance near vascular anastomosis. J Vasc Surg 1985; 2:419-423.

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CHAPTER 41 Bioinertness: An Outdated Principle David F. Williams

The Convention of Inertness

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iomaterials have been with us for the majority of the twentieth century. Their nature has evolved during this time, and the applications for which they have been used have increased in complexity and diversity. However, for much of this time, the functions required of these materials and the performance parameters of the medical devices in which they have been used have been relatively straightforward and largely confined to mechanical and physical characteristics. The selection and design of biomaterials have therefore been similarly constrained by the conventional engineering concepts underlying these applications. During the last few years the situation has changed quite radically as new concepts and new treatment modalities have required a fundamental reappraisal of the scientific principles upon which biomaterials science are based. Nowhere is this seen more vividly than in the use of biomaterials for implantable devices within the cardiovascular system, and specifically in relation to the application of tissue engineering concepts in reconstructive vascular surgery. In this chapter we attempt to provide a rationale for this fundamental change and to demonstrate the pivotal role of surface reactivity in the products and devices of the future. If we consider the historical role of biomaterials in implantable devices we can identify several clinical conditions that underlie their use, including congenital and developmental defects and trauma, but must recognize that it is in the treatment of diseases which have caused irreversible changes in tissues and organs that biomaterials play their biggest part. Thus, we see the widespread use of total joint replacements associated with the very widespread incidence of osteoarthritis, of intraocular lenses necessitated by cataracts, of prosthetic heart valves in the treatment of valvular disease, of dental implants following tooth and periodontal destruction, and of vascular prostheses and devices used for the prosthetic reconstruction of arteries. In all of these situations the functional requirements of the materials are simple and essentially nonbiological. The major implantable devices of the twentieth century either transmit loads, facilitate sliding motion, passively direct fluid flow, control fluid flow or simply fill a space. The materials that permit the devices to perform these functions need to have the appropriate mechanical properties and little else. Indeed, in the majority of these cases, the mechanical and physical requirements are not themselves very onerous, and a wide range of materials is theoretically available for the construction of these devices. In reality, of course, the situation is not quite so straightforward, since the devices usually have to perform their function for a long period of time within the confines of the human body. This implies that the materials have to supply these mechanical or physical characteristics without interference from the hostile physiological environment and without Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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causing any untoward or adverse effects in the patient. The latter two aspects are subsumed within the phenomenon of biocompatibility, and it has been a combination of the mechanical and physical properties together with biocompatibility considerations that have determined the selection of materials available for implantable medical devices at the present time. The term biocompatibility sounds simple to interpret since, as alluded to above, it implies compatibility or harmony with living systems. This concept, however, is too simple to be useful and the meaning of compatibility has to be explored further, both to understand the real logic of today’s material selection and the opportunities for the future. It is intuitively obvious that a biomaterial or an implanted medical device should cause no harm to the recipient. This is the underlying principle of biological safety, and it is this rather than the totality of biocompatibility which has shaped the historical evolution of biomaterials. Most crucially, the requirements of biomaterials have been dominated by the perceived necessity to be safe, which has usually been interpreted as a requirement that a biomaterial should be totally inert in the physiological environment and should itself exert no effect on that environment. In other words, there should be no interaction between biomaterials and their host, implying that the material should be nontoxic, nonirritant, nonallergenic, noncarcinogenic, nonthrombogenic and so on. This concept of biocompatibility which equates “biological performance” to inertness and biological indifference has resulted in the selection of a portfolio of acceptable or standard biomaterials that have widespread usage in applications such as those mentioned above. These range from the passive alloys such as stainless steel and titanium alloys, the noble metals of gold and platinum, some oxide ceramics such as alumina and zirconia, various forms of carbon and a range of putatively stable polymeric materials including silicone elastomers, polyolefins, fluorocarbon polymers and some acrylics and polyesters. The precise material chosen for an application from this type of list would depend on the precise mechanical or physical property specifications. It cannot be denied that this approach has led to the successful deployment of many types of prostheses in wideranging clinical areas, and, apart from a few notable and controversial situations, such implantable devices rarely impact adversely on the health of the patient. The concept of bioinertness has therefore served a useful purpose. Even when we look at these generally successful areas, however, it can be seen that there are significant limitations to their performance and to the type of patient in which they can be used with confidence. Total joint replacements are generally contra-indicated in the younger, more active patient; the performance of vascular prostheses decreases as the devices are used more distally; bioprosthetic heart valves give greater cause for concern in very young patients; and intraocular lenses may not give satisfactory results in patients with underlying inflammatory conditions. Questions may be raised, therefore, as to whether bioinertness is the most acceptable principle underpinning material selection or whether alternative concepts would be preferable.

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There are three fundamental reasons why bioinertness may not necessarily be the sensible choice. The first is that complete inertness is unachievable, so that any strategy predicated on this quality is ultimately doomed to fail. Secondly, if a device is made from materials which are inert and which do not interact with the body in any way, then it is unlikely that it can be truly incorporated into the body. For effective, long term performance in the dynamic tissue environment, it is usually preferable for there to be functional incorporation, which implies that the device should be stimulating the tissues to react to it positively rather than permitting them to ignore it. Thirdly, the emergence of tissue engineering principles, in which biomaterials are usually seen to act in a supporting role to biological components, has placed a very considerable emphasis on the utilization of intentionally degradable materials. These three features, which are determining the changing emphasis away from bioinertness, are explained below.

The Impossibility of Inertness The convention described above is based on the assumption that a prosthetic device replaces, augments or modifies tissues or organs and is intended to remain within the patient for the length of time that the function is required, typically for the natural lifetime of that patient. The inertness was considered to be of crucial importance, both to allow the device to retain its structural integrity and thereby maintain its performance, and to prevent or at least limit the release of any substance or component from the material which would have some toxicological or other adverse biological consequence. In the vast majority of circumstances, sufficient control has been exercised over the chemical stability of biomaterials to ensure that no catastrophic loss of that integrity occurs as a result of compromised bioinertness. There have been some exceptions and there are no grounds for undue complacency in this respect, but this is generally not the most important of considerations. On the other hand, the potential to initiate harmful biological processes in the host will always be finite as long as there are structures capable of degradation or components free to be released. Several features of the inherent structure of materials and of the complex nature of the tissue environment aggravate this situation. In particular, inertness in the physiological sense requires a great deal more than resisting degradation at the atomic or molecular level and, furthermore, even if it were that straightforward it would be extremely difficult to achieve. Indeed, it is now recognized that no material is totally inert in the body. Even those very stable materials mentioned earlier will interact to some extent with tissues. Titanium, although one of the most corrosion-resistant engineering alloys, corrodes in the body, judging by the presence of the metal in surrounding tissues. Gold and platinum will interact electrochemically with the saline-based extracellular fluid. For reasons not entirely understood, the most inert of the oxide ceramics suffer some long term changes within the body, and almost all known polymers will undergo oxidation, hydrolysis or other changes upon implantation.

Bioinertness: An Outdated Principle

With many materials, while the main component itself may be exceptionally inert, there are often minor components, perhaps impurities or additives, which can be released under some circumstances. The leaching of plasticizers and other additives from plastics provide good examples of interactions which are not necessarily related to the molecular breakdown of the material, but which nevertheless confer a degree of instability to the product. It should also be noted that descriptions of material degradation mechanisms have to take into account the special and indeed, unique, features of the tissue environment. Whatever its location, a biomaterial will continuously encounter an aqueous environment during its use. This is not simply a saline solution, however, but a complex solution containing a variety of anions and cations, a variety of large molecules, some of which are very reactive chemically, and a variety of cells. There are occasions when a degradation process can be explained, mechanistically and qualitatively, by the presence of electrolyte. This is the situation with most metals when they suffer from corrosion in a physiological environment. Even here, however, it is known that the kinetics of corrosion may be influenced by the organic species present, especially the proteins, and it is indeed possible for the corrosion mechanism to be somewhat different from that found in nonbiological situations. This phenomenon is even more pronounced with other groups of materials and it is clear that with polymers the kinetics and mechanism of degradation are fundamentally related to the precise details of the environment. Although hydrolysis remains the substantive mechanism for degradation of most heterochain polymers, including polyamides and polyesters, this hydrolysis may be profoundly influenced by the active species present in the tissue. In particular, the lysosomal enzymes synthesized and released from cells of the inflammatory response to biomaterials may influence the degradation process. Moreover, the hydrolysis may be supplemented by oxidative degradation, again occurring not only by virtue of passively dissolved oxygen in body fluids, but by active oxidative species such as superoxides, peroxides and free radicals generated by activated inflammatory cells such as macrophages. It is thus possible for homochain polymers not particularly susceptible to hydrolysis and not normally oxidized at room temperature to undergo oxidative degradation upon implantation. Polyolefins such as polyethylene and polypropylene come into this category. On the basis of a vast amount of experimental work and clinical experience, it is now clear that all biomaterials are inherently susceptible to some degradation process within the physiological environment. It is equally clear that although a variety of surface treatment methodologies are available to reduce or ameliorate the degradation, none of these are entirely effective and their availability does not negate the now accepted principle that complete inertness cannot be achieved. Moreover, in the context of interactions which affect the overall performance of the material in the physiological environment, it is important to note that an interfacial reaction involving a physicochemical process such as protein adsorption will inevitably take place, further emphasizing the fact that inertness is a very relative term and

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that there is no such thing as an inert biomaterial. It is never a question of whether a biomaterial will interact with the body but rather when and how. Under these circumstances, the preferred alternative strategy is to accept that such interactions take place and to attempt to control these interactions proactively and to incorporate these interactions into the design specifications.

The Requirement for Controlled Reactivity The above arguments indicate that complete inertness may not be possible, but still do not indicate that it is undesirable. The possibility that the products of any interaction between a biomaterial and its physiological environment could be released into the host has generally been considered a sufficient deterrent to utilizing any material that was significantly reactive in that environment. This logic is clearly only valid if those products were going to initiate an undesirable response from the host, either locally or systemically. If interactions took place whereby the products were totally harmless, then there would be less cause for concern over the inability to achieve inertness. More importantly, if the nature of the interaction were one which produced a more appropriate response from a tissue, then a positive virtue could be made of this lack of inertness. This quite different thinking is now enshrined in the current concepts and definition of biocompatibility. Instead of biocompatibility being equated with inertness, it is now recognized that it should encompass a wide range of reactivity, with the caveat that any reactivity is beneficial rather than harmful. On the basis of these ideas biocompatibility was redefined a few years ago as “the ability to perform with an appropriate host response in a specific application”. It is apparent that this definition still encompasses the situation where inertness is required, since the most appropriate response in some situations is indeed no response. A traditional bone fracture plate is most effective when it is attached mechanically to the bone and does not corrode. No response of the tissue to the material is required under these circumstances. In the type of device mentioned earlier in this chapter, in which the performance is dependent upon its physical replacement of diseased tissues and its incorporation into the structure of the body, inertness of all of the components may prevent optimal performance from being achieved. In particular, if a material is inert and unreactive within tissues, the long term host response will be associated with a lack of recognition and a lack of functional incorporation. An inert polymer such as polyethelyne or PTFE will induce the formation around it of a thin layer of collagenous fibrous tissue which can neither facilitate incorporation of the device into the tissue nor assist the device in achieving any of its functions. Moreover, this fibrous layer is unlikely to be stable and may alter its characteristics over time, this often being the ultimate cause of the device failure. In such circumstances, the biocompatibility characteristics of the materials in contact with their host tissues should be those which favor a positive interaction between the molecules of the material and the relevant molecules of that tissue, such that there is a functional attachment between the two. A total joint replacement which has the nonarticulating surfaces

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composed of materials or substances that are able to interact with the cellular components that promote bone regeneration at the interface should yield a more appropriate host response in the context of the long term mechanical stability of the joint. Similar arguments could be put forward for many of today’s implantable devices that are used in reconstructive surgery. This concept may be extended to many other types of implantable device, including those which function through an intervention in healing processes. For example, whilst it was noted before that conventional fracture plates need to be inert, on the basis that degradation or corrosion products could be harmful to the local tissue, it is not impossible for such devices to be made with surfaces that interact with osteoblasts, thus actively promoting or accelerating the healing process. Similarly, while many intravascular stents are fabricated from alloys chosen on the basis of their corrosion resistance and general inertness, the application of substances to the surfaces of these alloys which potentially have an ability to interact at the molecular level with those agents that are responsible for restenosis provides a new concept of bioreactivity rather than bioinertness for these devices. Almost every type of implantable device which has hitherto been associated with materials chosen for their apparent inertness may now be reconsidered in this light. This applies equally to devices where active surface molecules can control thrombogenicity, to devices which are able to encourage tissue repair and to devices which are able to minimize the undesirable consequences of inflammation.

The Essential Requirement of Reactivity in Tissue Engineering The desirability of some degree of reactivity in conventional implantable devices can be seen as a precursor to the essentiality of reactivity associated with the concepts and products of tissue engineering. It is important to remember here that the underlying principle of tissue engineering involves the combination of biological principles and substances with medical engineering principles and devices in order to provide products that are able to persuade the body to heal itself. In this rapidly emerging field, the major products utilize some material structure as a support, matrix or vehicle for the delivery of active biological molecules or cellular species to a target site in the host where repair or reconstruction is to be effected. Important as these material structures are, they do not normally constitute a component that is intended to remain in the body for very long, and certainly their function is incompatible with the characteristic of total inertness. There are several different levels at which reactivity is desirable or essential. As alluded to above, the most obvious of these is the desirability for intentional biodegradation in the polymeric support structures that form the basis of tissue-engineered reconstruction devices. Whether this is concerned with the repair of cartilage defects, of nerve tissue, of damaged skin or diseased tissue within the vascular system, the template of a device usually involves a polymer that can degrade over a period of time ranging from a few weeks to a few years and where the degradation profile is consistent

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with the delivery of active tissue repair promoting molecules and with the physical support necessary for the developing tissue structure. It goes without saying of course that such a material should not degrade to produce harmful side effects, which could be associated with either the chemical characteristics of any degradation products or their physical morphology. Although this is not a trivial matter, such that inflammatory reactions associated with the degradation cannot be ruled out, there are several biodegradable polymers which may now be used with reasonable confidence in these situations. Tissue engineering products should, however, involve more than a degradable structural polymer. The real essence of such a product is the biological activity which it imparts to the host site. A simple, empty, porous scaffold of a degradable polymer implanted within connective tissue may become infiltrated with repair tissue and may ultimately degrade to leave an area of reconstituted tissue derived from that infiltration. This, however, is not tissue engineering, since the tissue will not have been directed or controlled in terms of its structure and nature. In order to be really useful, this product would have to incorporate the appropriate cellular components, normally previously derived from the hosts, or certain molecules that are able to signal to the host cells the appropriate instructions to produce the desired tissue morphology and function. Under these circumstances, the biodegradable polymer has to be chosen very carefully, such that at the very least it will be compatible with the desired cell function in its vicinity or, preferably, such that it is capable of delivering these cell signaling molecules with the desired activity and in the desired manner. Inherent inertness is therefore unacceptable in these tissue-engineered products. The challenge is to introduce the desired level of reactivity without compromising the biological safety of the material.

Conclusions For all the reasons outlined above, bioinertness can be considered as an outdated principle. Although there are some types of implantable device which are still better served by inert materials, the majority should preferably incorporate intentionally interactive materials, whilst others, especially in this new area of tissue engineering, positively demand this characteristic. It has to be said, however, that academic arguments in favor of replacing a concept that is perceived to be out of date with a new principle cannot be sustained unless this new principle can be verified and reduced to practice. In view of the fact that bioinertness has never produced any harm, it is indeed a brave new world which is predicated on the availability of biomaterials that are both functionally interactive with their host and intrinsically safe. Many might argue that these in fact are contradictory requirements. Clearly that need not be the case, but the whole future of tissue engineering products that are able to demonstrate efficacy and safety is dependent on the development of the appropriate materials. However outdated bioinertness may be, it should not be replaced by materials based upon hope rather than scientific reality.

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CHAPTER 42 Bioinert Biomaterials: Are Their Properties Irreplaceable? Patrick T. Cahalan n 1955 Sewell and colleagues1 performed a study comparing ovine and bovine sources of catgut sutures in three animal models. The objective of the study was to quantitatively compare the tissue response to the implanted material, and this event is often cited as the beginning of biocompatibility testing. Coincident to this was the rapidly expanding postwar technology in high performance plastics, with mounting evidence that biological grafts, even homografts, were not going to survive the challenges in vascular graft applications.2 Voorhees’ observations in 1952 of a neointima free of thrombi on a silk thread, and his later experiments with Vinyon “N” cloth in dog aortas, clearly showed that a synthetic material could serve as a conduit, and be accepted by the tissue.3 Studies would follow showing the merits of other materials, particularly Teflon and Dacron.4 The paradigm was set, and materials were classified as very reactive to inert. What followed was once described as “20 years of frustration”.5 During this period commercial efforts were driven by the promise of a very large market, estimated between 0.6 to 1 billion dollars annually. Scientific efforts were focused on comparing properties of materials with limited physiological markers, primarily the occurrence of thrombi on the material surface. An example of surface property concepts that were hypothesized to be beneficial are seen in Figure 42.1.6 The majority of the surfaces listed are in essence attempts to create surfaces with minimal effect, whether it be termed low protein or platelet adhesion. In the same 1972 article, Baier refers to much of the individual surface property characterization towards blood compatibility as people working in a maze within a maze. Hydrophilic surfaces, for example, while showing low platelet adhesion and therefore thought to be blood compatible, when tried in a vascular graft position did not prove to provide patency. Thus, researchers thinking they had found a way out of the blood compatibility maze found themselves in another maze with new barriers to solve. Baier’s representation of the surface properties compared to the living blood vessel intima serves to highlight the focus of blood compatibility to the vascular application. This narrow focus, plus the lack of clearer understanding of cellular and molecular biology, helped continue the quest for the “holy grail” of inert materials. Not until it was shown in the late 1970s that endothelial cells were more active in maintaining blood compatibility7 did the paradigm begin to change. Until this time it was even possible to hypothesize that endothelial cells were the perfect inert surface. The emergence of RIA and ELISA in the mid-70s opened up the possibilities of looking at the molecular and cellular aspects involved at the surface. The common use of these tools did not make its way into blood compatibility research until well into the 1980s, and today a significant number of papers are still offered that rely largely on platelet adhesion alone to determine blood compatibility of surfaces. The increase in understanding of reactions between coagulation factors and formed elements of the blood has led to a change in thinking that

I

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Fig. 42.1. General approaches to improve blood compatibility: early years (70s).

previously focused on the bulk reactions in blood to the realization that these critical reactions “occur almost exclusively on surfaces”.8 The formerly thought to be inert endothelial cell has been shown to have as many procoagulant properties as anticoagulant.9 The pat answer for years to the question, “What is the perfect surface?” was: “An endothelialized surface.” The fact that an endothelialized surface can be thrombogenic opens up a new maze. It also begs the question, “Bioinert biomaterials: Are their properties irreplaceable?” “The long sought after biologically inert material has not as yet been developed, and the realization of the inherent ‘biointeractivity’ of all implanted polymers has led many investigators to pursue strategies aimed at optimizing the biological interactions with the synthetic polymer rather than to minimize all biological interaction.”10 This statement would seem to put an end to the question, “Is there a need for bioinert materials?” Unfortunately, vascular grafts and other biomedical devices still use synthetic materials of construction, and their replacement with other biological materials or repair therapies is not on the very near horizon. The great promise of biodegradable materials has not been realized, and attempts at endothelialization of synthetic materials, while showing promise, have not consistently exceeded the performance of autologous tissue. Preadsorption of synthetic vascular grafts with cellular adhesive proteins tends to enhance the short term attachment of endothelial cells, but platelet adhesion, and therefore increased thrombogenicity, typically increases with time.11 A possible

explanation for this failure may be the nondiscriminatory nature of some of the cell adhesion receptors to adhesion proteins. Some new approaches to deal with the nondiscriminatory cell receptor binding to adhesion proteins attempt to attach peptide adhesion ligands to surfaces.12 A critical requirement stated by Hubbell in such attachment is to render the surfaces resistant to protein adsorption by first grafting with a hydrophilic nonionic polyethylene oxide (PEO) chain to which the peptide ligand could be tethered. The latter stated requirement could serve to finally propose a more pragmatic application of the “inert” concept. Inertness with respect to a nonionic hydrophilic spacer that is resistant to protein adhesion permits a selectivity or control of molecular and cellular events at surfaces. The challenges left with this approach are to achieve the proper density of ligands on the surface, their orientation, and the ability to accomplish this on synthetic substrates in the geometric configuration of a biomedical device such as a vascular graft, artificial heart, heart valve, or ventricular assist device. While this elegant concept represents a major advance in the effort to make an endothelialized and hopefully long term blood-compatible surface on a vascular graft, it is not certain whether another maze will not open up, such as continued hyperplasia. Due to noncompliance of synthetic structures, some might argue that a truly new totally functioning vascular tissue might not be possible, and that even though there is an endothelial layer on the lumen, the structure will succumb to other mechanisms of failure. More optimistically, it could be hoped that compliant structures

Bioinert Biomaterials: Are Their Properties Irreplaceable?

could be engineered that also could be modified to promote spatially correct cellularization of smooth muscle cells, and even include neovascularization. The progress made by persistent researchers10,12 in the vascular graft field suggests that the latter possibility is not out of the question. As mentioned earlier, the vascular graft focus may have actually been in part responsible for a lack of more rapid progress in understanding the cellular and molecular events at the surface, particularly in development of more bloodcompatible surfaces. Other application areas such as dialysis and cardiopulmonary bypass (CPB) have created a strong research effort in providing surfaces with enhanced blood compatibility, and in this effort industry, academia, and clinicians have invested large amounts of time and capital in investigating improved biocompatible surfaces on devices not necessarily designed for permanent implant. From a strictly clinical perspective, the ability to monitor coagulation and inflammatory responses to foreign surfaces during CPB, and evaluate devices after use, has created a plethora of scientific publications involving measurement of molecular and cellular events on blood-contacting surfaces. This has also created a commercial stimulus for funding research on blood materials interactions and research in surface modification of biomedical devices. In our laboratory, surfaces deemed “inert” have been studied using multiple ELISA markers such as TAT, F1.2, elastase, and SC5b-9 (terminal complement complex) in addition to measuring platelet adhesion and activation. The most often cited inert molecule is poly(ethylene glycol) or (PEG). It has been suggested by some that a PEG surface has little or no protein adsorption or interaction.13,14 This premise has also been challenged by some researchers.15,16 It has further been suggested that PEG surfaces may serve to delay rather than reduce protein adsorption and activation.17 With respect to blood compatibility, we found that while a PEG surface does in fact show lower protein adsorption and platelet adhesion, it does not show any improvement in TAT, complement (SC5b-9), or elastase. These studies used a blood loop system that contained a valve to effect physiological flow, and human blood containing 1 U/ml of heparin. While PEG surfaces have been studied and widely reported to approach inertness, it is obvious that there is yet much to be learned before this technology is optimized. Nevertheless, the suggested use of this molecule to tether a peptide, or possibly a specific adhesion molecule, may in the end prove to be the best application of PEG. Another molecule that is receiving considerable attention for its low protein and platelet adhesion is phosphatidylcholine (PC). Methacrylate copolymers with pendant PC groups (MPC) have been coated with a facile process to polymers, and the PC moiety is expressed to a high degree on the uppermost surface.18 Conceptually, the MPC coating is said to mimic the natural phosphatidylcholine present on the surface of nonactivated platelets. In testing in our laboratory, this surface does show reduced platelet adhesion and less TAT formation. Additional testing needs to be completed to evaluate the complement and white cell response to this surface. In our hands this polymer has shown material-dependent coating characteristics, and evaluation is tedious,

465

as optimal pretreatments followed by confirming XPS and TofSIMS data is required to assure optimal presentation of PC moieties on the surface before testing. There is supporting evidence for the inertness of a pure PC surface.19 Studies performed at the University of Maastricht have demonstrated that a phospholipid bilayer of pure dioleoyl-phosphatidylcholine (DOPC) on an ellipsometry plate is indeed an inert surface, and assembly of the prothrombinase complex requires the addition of dioleoyl-phosphatidylserine (DOPS) in significant amounts.28 DOPS is present on the internal surface of the platelet membrane, and platelet activation causes a flip-flop of the DOPS to the external surface of the membrane. When DOPS is on the surface, the prothrombinase complex can be assembled and thrombin generation can take place. Unfortunately, the stability of a DOPC bilayer is very poor, and at this time has not been attained with consistency on a polymeric surface. The MPC polymer approach would be to assure uniformity and density of the PC moiety on the surface, and thus impart through mimicry an inertness to the surface. While the DOPC surface appears to be the closest thing to an inert surface, the actual surface of the plasma membrane is in fact made of more than just DOPC. Transmembrane proteins assure that platelets in the end are very reactive with their environment. It is perhaps appropriate to reflect back to the comments of Spaet8 with respect to all important reactions in the body taking place on surfaces, and that these surfaces we have learned to be anything but inert. In the end, one of the roles or properties of inert biomaterials that is needed is their ability to enhance the expression of another molecule, such as a peptide as suggested by Hubbell. This may be accomplished by minimizing competing phenomena such as protein adsorption, as well as presenting a molecule in a conformation more accessible to its target receptor. This property is indeed possibly irreplaceable. Our studies specific to the attachment of heparin to surfaces clearly indicates that the bioactivity of an attached molecule is dependent on its ability to interact with target molecules, and the ability to control the loading and the reactivity of the attached molecules is critical to achieve optimal functioning on the surface.20 Heparin on these surfaces was tethered to a nonionic hydrogel that was derivatized to achieve different loadings of heparin. Here the surface is certainly not inert, but it is designed to optimally compete with the undesirable physiological mechanisms of coagulation and platelet adhesion. These surfaces do not prevent protein adhesion, but the proteins that have adhered show different relative concentrations to control surfaces, and the performance of the surfaces correlates with these adhesion profiles.21 In nature, a denuded endothelial lining exposes a procoagulant surface, yet this surface can rapidly passivate without excessive thrombin formation. Fibrinolysis, controlled inflammation and, finally, remodeling give a normal healing response that restores the tissue to a functional state, barring the introduction of complications of a biomaterial, ongoing disease state or flow restrictions. A normal healing response replaces inertness with spatially correct signals via the subendothelial extracellular matrix and thus controls the first step of repair, which is platelet

466

adhesion and coagulation. A naturally formed and controlled hemostatic response (clot) presumably is resolved without the excessive inflammatory reactions associated with the presence of foreign materials. If those foreign materials possess an inert surface, the resultant histology should resemble that of a injury without the presence of a foreign material. Since injury exposes the blood to nonendothelialized surfaces, and these surfaces do bind plasma proteins, it might be a better approach to attempt to modify a surface to have a similar protein adsorption profile. Implants of biomaterial surfaces modified with collagen (data to be published) after 10 weeks show minimal capsule formation with microvascular structures within two fibroblast layers of the biomaterial surface. The control biomaterial surface (polyurethane) has macrophage and foreign body giant cells (two layers) followed by 12-15 layers of densely packed fibroblasts adjacent to the material surface (a typical fibrotic response). The attachment of the collagen to the above mentioned surface was not simple adsorption, but rather a covalent attachment with isolation of the collagen layer to the uppermost layer (70 Å).22 XPS data showed a pure surface of collagen, and this was attainable based on the inherent properties of the grafted surface beneath. This should be a target for modification of surfaces whether inert or interactive, and that being complete and uniform coverage without influence from intermediate chemistries used to couple the molecules to the surface. In addition to the property of inertness being used to enhance the presentation of molecular species to the physiological environment, there is another possibly irreplaceable property of inertness. That is the relative chemical inertness of biomaterials. In addition to finding the optimal protein adhesion mosaic on a surface, the biomaterial of construction of a device must withstand the degradative environment of the body. Numerous attempts have been made, for example, to fabricate vascular prostheses using polyurethane, when several instances of urethane degradation in vivo have been documented.23,24 Attempts to mimic the in vivo response in an in vitro system have shown that adsorption of the plasma protein α2-macroglobulin could play a role in biodegradation of the urethane.25 It might be hypothesized that a surface modification that could isolate the biomaterial surface from protein and cell attachment could serve to protect its stability. Those who would choose to continue the use of polyurethanes for vascular grafts will need to deal with the potential degradation mechanisms demonstrated. An alternative approach to elastomeric materials may be fabricated grafts using very stable materials such as PTFE and PET that have mechanical properties created by the bulk design of the device. Biodegradable materials have boasted the promise of ultimate inertness in that they are not left behind for reaction with the body. While very impressive advances are being made in resorbable polymers,26 their application at this time may still be limited by the need for mechanical properties that are best met with traditional biomaterials. An additional issue with resorbable materials is the control of the rate and type of resorption, ablative or eroding. It is also somewhat taken for granted that the inflammatory responses

Tissue Engineering of Prosthetic Vascular Grafts

of these materials can easily be controlled by coupling the release of anti-inflammatory drugs; this may possibly open up yet another maze. In summary, the 20 years of frustration27 to which Andrade referred have been replaced by numerous multidisciplined efforts, aided by a much better understanding of molecular and cellular events at surfaces. The knowledge gained has opened up many approaches to finding the holy grail, and also created a number of positive spin-offs in therapies apart from vascular grafts. The concept of bioinert materials still persists, not as an end in itself, but rather as a tool to enhance positive mechanisms or to retard undesirable competing mechanisms. In the quest for new biomaterials, and for tissue engineering to replace diseased tissue with new functioning tissue, we will surely progress in an iterative manner. Mundane issues such as sterilization may render new concepts nonfunctional, but may stimulate new sterilization technology in order to bring them to market. Mechanical property limitations may require creative processing or fabrication advances. Promising surface modifications may first produce composite hybrid structures with traditional materials that can later be replaced by resorbable materials. What in the end is probably irreplaceable is the learning process involved in trying to find inert biomaterials. References 1. Sewell WR, Wiland J, Craver BN. A new method of comparing suture of ovine catgut with sutures of bovine catgut in three species. Surg Gynecol Obstet 1955; 100:483. 2. Szilagyi DD, McDonald RT, Smith RF et al. Biological fate of human arterial homografts. Arch Surg 1957; 75:506. 3. Voorhees AB, Jaretski A, Blakemore AH. The use of tubes constructed from Vinyon “N” cloth in bridging arterial defects. Ann Surg 1952; 135:332. 4. Deterling RA, Bhonslay SB. An evaluation of synthetic materials and fabrics suitable for blood vessel replacement. Surgery 1955; 38:71. 5. Andrade JD, Nagaoka S, Cooper S, Okano T, Kim SW. Surfaces and blood compatibility: Current hypotheses. Vol XXXIII Trans Am Soc Artif Organs 1987. 6. Baier RE. The role of surface energy in thrombogenesis, Bulletin of The NY Academy of Medicine 1972; 48(2):257-272. 7. Weskler BB, Marcus AJ, Jaffe EA. Synthesis of prostaglandin 12 (prostacyclin) by cultured human and bovine endothelial cells. Proc Natl Acad Sci USA 1977; 74: S.3922-3926. 8. Spaet TH. Blood in contact with artificial surfaces: Where have we been and where are we going. Annals of The New York Academy of Sciences, 1987; 516. 9. Gimbrone Jr MA. Vascular endothelium: Nature’s blood container. Vascular Endothelium in Hemostasis and Thrombosis. Gimbrone Jr MA, Ed. Edinburgh: Churchill Livingstone, 1986; 1-13. 10. Greisler HP, Gosselin C, Ren D, Kang SS, Kim DU. Biointeractive polymers and tissue engineered blood vessels. Biomaterials 1996; 17:329-336. 11. Seeger JM, Klingman N. Improved endothelial cell seeding with cultured cells in fribronection-coated grafts. J Surg Res 1985; 38:641-647. 12. Massia SP, Hubbell JA. Tissue engineering in the vascular graft. Cytotechnology 1992; 10:189-204.

Bioinert Biomaterials: Are Their Properties Irreplaceable? 13. Merrill EW, Salzman EW. Polyethylene oxide as a biomaterial. Am Soc Artif Intern Organs 1993; 6:60. 14. Harris JM. Poly(ethylene glycol) Chemistry: Biotechnical and Biomedical Applications. New York: Plenum Press, 1992:15. 15. Llanos G, Sefton MV. J Biomater Sci Polymer Edn, 1993; 4:381-400. 16. Sheth SR, Leckband D. Measurements of attractive forces between proteins and end-grafted poly(ethylene glycol) chains. 17. Kulik E et al. Poly(ethylene glycurface: Reduced vs delayed protein adsorption and activation. Abstract from Firth World Biomaterials Congress, Toronto, Canada. 1996; 29 May-2 June. 18. Ishihar K, Nakabayashi N. Part A: Polymer Chemistry: Specific interaction between water-soluble phospholipid polymer and liposome. J Poly Sci 1991; 29:831-835. 19. Andree HAM. Phospholipid binding and anticoagulant action of annexin V. Ph.D. thesis at the University of Maastricht, The Netherlands 1992. 20. Lindhout T et al. Antithrombin activity of surface-bound heparin studied under flow conditions. J Biomed Mater Res 1995; 29:1255-1266. 21. Sapatnekar et al. Blood-biomaterial interactin in a flow system in the presence of bacteria: Effect of protein adsorption. J Biomed Mater Res, 1995; 29:247-256. 22. HendriksM. A study on the covalent surface-immobilization of collagen. Ph.D. thesis, Development of biomaterials

467 with enhanced infection resistance. University of Eindhoven, The Netherlands. 1996:127-148. 23. Stokes K. Environmental stress cracking in implanted polyether-polyurethanes. Polyurethanes in Biomedical Engineering. Plank H, Egbeers G, Syre I, eds. Amsterdam: Elsevier, 1984:243. 24. Takahara A, Hergenrother RW, Coury AJ, Cooper SL. Effect of soft segment chemistry on the biostability of segmented polyurethanes. II. In vitro hydrolytic degradation and lipid adsorption. J Biomed Mater Res 1992; 26:801-818. 25. Schubert MA, Wiggins MJ, Schaefer MP, Hiltner A, Anderson J. Oxidative biodegradation mechanisms of biaxially strained poly(etherurethane urea) elastomers. J Biomed Mater Res 1995; 29:337-347. 26. Sawhney AS, Pathak CP, Hubbell JA. Bioerodible hydrogels based on photopolymerized poly(thylene glycol)-co-poly(α-hydroxy acid) diacrylate macromers. Macromolecules 1993; 26:581-587. 27. Andrade JD, Coleman DL, Didisheim P, Hanson SR, Mason R, Merrill E. Blood-materials interactions: 20 years of frustration. Trans Am Soc Artif Intern Organs 1981; 27:659-662. 28. Smeets E. Scrambling of Membrane Phospholipids in Platelets and Erythrocytes. Ph.D. Thesis, University of Maastricht, The Netherlands, 1996.

Scaffold Engineering Material Aspects

CHAPTER 43 Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses Arthur J. Coury

Introduction

T

he successful implementation of any medical device requires a systematic development process from concept through use in humans. Rigorous quality systems and design controls are now mandated by law,1 and the framework they provide2 is especially relevant to the development of systems as complex as tissue engineered vascular prostheses. The selection of polymeric materials to serve as structural scaffolds for the vascular prostheses should only be made in the context of the design intent of the device. In the spirit of design control, this chapter begins with a concept statement—the first stage of the design process. The intent is not to completely design the device but to provide a minimum set of characteristics which will set rather loose boundaries, to allow a range of prosthetic materials to be considered in this chapter.

Design Concept The product is a vascular prosthesis intended to address the unmet medical need for functional replacement of small (≤ 6 mm) and medium (6-12 mm) diameter arterial segments.3-4 The device consists of a structural matrix (scaffold) of available synthetic biomaterials and possibly other components assembled prior to implantation. The scaffold promotes and retains attachment and/or ingrowth of biological tissues. The synthetic scaffold is biostable, that is, resistant to excessive structural deterioration for the projected lifetime of the device, which is considered to be a “permanent” implant. The synthetic components of the device elicit an acceptable host response alone or through pharmacologic intervention. The device can be produced by available techniques into configurations which provide adequate hemodynamic flow. It may be packaged, sterilized, stored and implanted by accepted processes. It is amenable to any special techniques (e.g., in vitro cell seeding or surface modification) which are required for its proper function.

Material Requirements and Limitations Given the design intent of the device, specific material requirements for the structural matrix are discussed as subtopics of this section. It must be stated at the outset that no material is optimal in all of the requirements. The final selection of a material will involve compromises from the ideal, but demonstration of acceptable characteristics by appropriate testing is required.

Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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Availability by Purchase or Synthesis Particular attention should be paid to long term availability of the polymer. In recent years, major suppliers of polymers have withdrawn their products from consideration for implantable devices, 5-6 chiefly for liability reasons. In the United States, several legislative initiatives have led to the passage of the Biomaterials Access Assurance Act on July 30, 1998 (Public Law 105-230) which offers protection to biomaterials suppliers. The effect of this legislation will likely take several years to be fully realized. The supply problem is especially serious for stable biomaterials which often require capital-intensive processes available only to large chemical companies—the sources of highest liability exposure. This problem has led device manufacturers to search internationally for biomaterial equivalents or to work with smaller companies with limited manufacturing capability, but lower potential liability consequences. Fabricable into Intended Configuration Reasonable mechanical characteristics of the tubular device include flexural compliance, kink resistance and suturability. While materials with a wide range of physical properties can achieve these goals, there are propertydependent limitations on the configurations of these materials that can be used. Fiber-based or expanded wall configurations are probably required if the material has a relatively high to medium modulus (e.g., poly(ethylene terephthalate), polypropylene, polytetrafluoroethylene). Low modulus compositions such as elastomers (e.g., polyurethanes) can be considered as solid-wall, porous-wall or fiber-based tubes. Extrusion, molding or casting processes must be available to produce the base configuration. If surface modification is required, acceptable processes must be applicable to the material. Stable to Processing, Storage and Use Conditions Maintenance of structural integrity by the biomaterial scaffold is the main requirement of this chapter and is addressed in the sections on degradation mechanisms and discussions of individual materials. The major point to make here is that there are stability considerations at each of the many stages of processing and storage of the biomaterial, not just during residence in vivo.8 The material, at the time of implant, is the product of its prior history and attention to preserving its integrity throughout can assure stability of the implant. Elicits Acceptable Host Response for Life of Device Inherent nonthrombogenicity of the biomaterial is not likely and is not assumed to be required. A biological response will be induced, and control of that response is the function of the tissue engineering component of the vascular product. It is assumed that the matrix material has passed a standard biocompatibility protocol2 and that control of other responses can be achieved by mechanical design, local or systemic drug delivery, surface or bulk modification. Any additives used for processing, stabilizing or effecting biological responses are likewise assumed to be nontoxic.

Tissue Engineering of Prosthetic Vascular Grafts

Modes of Polymer Degradation The requirement of biostability does not imply that the biomaterial is inert in the medium under consideration. It must be appreciated that a polymer, and all materials, respond continuously to changes in environment before and during the implantation period.8 All polymers are permeable to gases and liquids. They absorb, transmit and desorb components. Their surface and bulk properties change with time and the imposition of external forces, sometimes reversibly, other times irreversibly. The changes can be favorable or unfavorable.8 Degradation leading to compromised performance of a scaffold in vivo can be physical, chemical, or a combination of both. Both types of degradation can also necessitate rejection of a component or device prior to use.9 Physical changes may include swelling, plasticization, crystallization, decrystallization, fatigue fracture, creep, simple stress cracking and kinking, among others.10 Chemical degradation generally consists of covalent bond cleavage, but, in some cases, covalent crosslinking or ionic bond transformations are involved. Variables such as heat, moisture, oxygen, light, radiation, even mechanical stress, induce chemical degradation.10 Many or all of these factors may be encountered by a biomaterial prior to its implantation. Each of these destructive agents is generally manageable to acceptable levels by process and exposure control and the use of additives10 if the material is inherently stable in the body. The longest intended phase in the lifetime of a “permanent” implant is residence in the body. The polymeric scaffold for a tissue engineered vascular prosthesis is subject to all of the forces listed above which may induce physical degradation. Appropriate device design and processing (e.g., to provide burst strength and minimize residual stress) and proper implantation techniques (e.g., to minimize kinking and applied stress) must be implemented to minimize physical degradation.10 Chemical degradation forces are also encountered at every phase of the biomaterial scaffold’s lifetime. The objective is to understand these forces and the material’s susceptibility to them and to protect against their effects. For example, antioxidants can protect polymers such as polypropylene and polyurethanes from thermooxidative, photooxidative and autooxidative (oxygen induced) degradation.10 In the body environment, modes of chemical degradation consist of three major types: hydrolysis, oxidation and mineralization (calcification). The latter mode may or may not induce significant covalent bond cleavage, but does involve chemical transformations that cause structural damage to the polymer and will be treated as one of the three subtopics discussed next. Hydrolysis Hydrolysis is the cleavage of susceptible chemical bonds with water. Simple hydrolysis can be catalyzed by acid or base and occurs as a consequence of the number and reactivity of the cleavable bonds. Enzyme-catalyzed hydrolysis (e.g., by esterases, proteases) has been reported9,11-14 but, for most synthetic polymers, enzymatic effects are superficial and minor and will not be further discussed here.

Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses

Hydrolytically susceptible bonds generally consist of carbonyl groups with a heteroatom (O,N,S) on one or both sides of it. These structures and several other hydrolyzable groups are listed in Figure 43.1. Hydrolysis rates differ among the susceptible groups. For example, carbonyl group reactivity decreases in the order: anhydride > ester > urethane > amide. Ethers, ketals and acetals are labile to acidic pH but stable to basic conditions. Backbone structure and intermolecular interactions of hydrolyzable polymers also affect degradation rate. Factors such as crystallinity, hydrophobicity and crosslink density tend to decrease hydrolysis rate9 and render polymers based on hydrolytically labile groups potentially viable candidates for long term implant. Polymer structures that resist hydrolysis include hydrocarbons, silicones, sulfones, halocarbons and isolated carbonyl-containing molecules (i.e., ketones).

Inside the body, biomaterials are subject to oxidative insult as a consequence of inflammation due to the foreign body reaction (phagocytic attack).9,17 Powerful oxidative products result from phagocyte activation (e.g., hydroxyl radical, peroxynitrite, hypochlorite), which can attack oxidizable bonds. Susceptible sites generally are those that can stabilize a free radical (Figure 43.2).9 Oxidatively stable structures include straight-chain hydrocarbons, halocarbons and fully-oxidized groups (e.g., ketones, sulfones, carbonates, esters, urethanes). We have, however, produced some evidence that the carbon atom next to the heteroatom of the latter carbonyl groups may be amenable to oxidation.17 Evidence for direct (receptor-ligand) enzymatic catalysis of oxidation is limited and inidcative of minor consequences.12,18 In vivo oxidation of polymers has been most extensively studied for polyurethanes and polyethylenes.9,15-24 Structural degradation due to oxidation is usually manifest as surface fissuring, deep crack formation, fragmentation or wear, often in zones of applied stress. Radiation sterilization (e.g., of polyethylene) can produce long-lived radicals with autooxidation and embrittlement leading to increased wear and fragmentation in use. Mitigation of structural degradation due to oxidation can be accomplished by minimizing residual and applied stress, controlling exposure to destructive radiation, use of antioxidants, isolating susceptible polymers in a device from direct attack by phagocytes or soluble oxidants and, finally, by the use of oxidation resistant materials.9,10,15,16

Oxidation Oxidative degradation of polymers involves destruction or modification of molecular structure via electron transfer reactions.9 The polymers to be considered have varying degrees of susceptibility to oxidation. Prior to implantation, specific operations such as melt processing, radiation sterilization or exposure to light containing ultraviolet wavelengths can promote oxidative degradation. Autooxidation involves reaction with molecular oxygen and is a likely mechanism of many oxidative events.15 Other oxidants (e.g., metal ions, hypochlorite, hydroperoxides)16 are also effective in causing degradation.

O X C Y

O

O

C Z

C

Fig. 43.1. Hydrolytically susceptible groups.

X=O,N,S,H Z=O,N,S Y=O,N,S,H X or Y≠H Hydrolyzable carbonyl structures

O CH2 C N (nitrile)

(ether) O CH2 O

(ketal or acetal)

O (cyanoacrylate) O P O (phosphate) C N C OR O O O S NH (sulfonamide) R=alkyl O Other hydrolyzable groups CH

CH

471

472

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 43.2. Oxidatively susceptible groups.

*

* * CH

CH

X

R hydrocarbon chain branch (R=alkyl, aryl, allyl)

X=O,NH1-2 ,S (eg, ether, amine, alcohol, sulfide)

*

OH

*C

O

H aldehyde

phenol * = susceptible site

Calcification Calcification or mineralization involves the deposition of calcium phosphate salts on or within implanted structures. These phenomena are termed extrinsic and intrinsic calcification respectively. Extent of calcification is related to the level of soluble calcium and the device function, more than the chemical composition of the material.25,26 Calcification has been reported to occur with synthetic vascular prostheses, causing stiffening and obstruction.25 The most prominent effects of calcification occur, however, in devices that undergo extensive cyclical deformations in vivo (e.g., prosthetic heart valves, heart assist devices).25 Generation of microscopic defects in the polymer may produce nucleation sites for the initiation of crystal formation, which can lead to extensive mineralization and, ultimately, mechanical failure of the device. Chemical degradation of the biomaterials has not yet been implicated in initiating or promoting calcification. Anticalcification additives have been applied to synthetic biomaterials with some success.26 Vascular implants that undergo moderate deformation should present a relatively low risk for calcification.

Polymers with Potential to Serve as Biostable Scaffolds This section describes polymers with potential for use as scaffolds for permanent tissue engineered vascular prostheses. Commercial polymers used in existing prostheses will be described first. Polymers that may not have been commercialized but have been described in the literature for use in medical devices will be briefly described. Finally, commercial polymers that have not been reported in vascular prosthesis literature, but may satisfy the design requirements for scaffolds listed above, will be mentioned. All of the polymers to be described are assumed to contain additives such as processing aids and stabilizers. Different grades, lots or brands of polymers with the same generic structure may be more or less stable as a result of additives, processing conditions, molecular weight differences, crystallinity differences and other variables.

The emphasis of this section is primarily on biostability, although product design and biocompatibility are not ignored. The general assumption is that the latter two requirements can be met by approaches described in this book if the scaffold can perform its function for the intended period. O

O

Poly(Ethylene Terephthalate)(PET) PET vascular prostheses have dominated the large diameter field since their introduction in 1957.27 PET displays several advantages in its “standard” or “micro-denier” fiber form. It is a strong material having a tensile strength in oriented form of 170-180 MPa and a tensile modulus of about 14,000 MPa.28 It is readily fabricable into woven or knitted textile or mesh form. It can be crimped to enhance kink resistance. It is relatively stable. Although it is based on the hydrolytically susceptible ester group, it is highly crystalline (MP ~ 265°C) because of its ordered chain structure and deep drawing during fiber formation.9 Implant durability studies have shown or estimated progressive deterioration of physical properties over decades as a consequence of hydrolysis, and, possibly, some oxidation induced by activated phagocytes.29-31 In vivo studies have projected approximately 30 years to full resorption in humans.29,30 PET is relatively stable to sterilization by ethylene oxide and ionizing radiation;32,33 however, techniques involving steam, dry heat or chemically reactive sterilants32-34 may cause deformation or degradation. Mitigating against the use of PET for small diameter prostheses is the high reactivity of blood and vascular tissue toward the implant with consequent high inflammation, neointimal proliferation and inhibition of cellular regeneration. This response to PET has been demonstrated with textile vascular prostheses35-39 and vascular stents in open mesh configuration.40,41 The blood and vascular tissue reactivity of PET can vary with the proprietary additives and treatments applied to the individual manufactured devices.42 PET has successfully served as the structural matrix for composite vascular devices. Albumin, collagen or gelaOC

COCH2CH2 n

Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses

tin impregnation of PET textile prostheses is performed to seal them against blood leakage.38-43 Healing responses of the composite are comparable to the unsealed PET.44 The Omniflow Vascular Prosthesis (Bio Nova, International) is formed from PET mesh placed on a silicone mandrel and implanted in sheep to produce a collagen-PET composite device. It has been used successfully for peripheral vascular replacement.45 Other modifications, such as incorporation of bioactive molecules46,47 and surface modification for blood compatibility48 have shown promise. In sum, PET should be considered a highly credible candidate for biostable vascular scaffolds. Polytetrafluoroethylene(PTFE) CF2CF2 n PTFE is a member of the fluorocarbon class of polymers. It is, by far, the most commonly used fluorocarbon49 in implants because of its demonstration of long term biostability and biocompatibility in vivo. PTFE vascular prostheses are prominent in the medium diameter (~7-9 mm) and dialysis access markets.27 The polymer is prepared in powdered form from tetrafluoroethylene monomer. The product is a highly crystalline material (> 90% crystallinity) which is not completely fusible, but is formed into shapes by sintering.49 A densely sintered configuration approaches the intrinsic properties of the material and gives a moderate stiffness (tensile modulus of elasticity = 0.5 GPa) and tensile strength (14 MPa).49 Implantable forms of the material of interest for vascular prostheses include the “expanded” and textile products.50 Expanded PTFE (ePTFE) is made by an extrusion, drawing and sintering process50 to produce a tube with a porous wall consisting of fibrils and nodules which is controllable to different pore sizes (e.g., 30 and 60 µm). The porosity can be used advantageously to promote ingrowth of tissue and formation and retention of an endothelial layer in vascular prostheses.51,52 Knitted PTFE cloth is effectively used as sewing rings for heart valve prostheses,53,54 and should be considered when a textile configuration for a biostable scaffold is preferred. It should be noted here that the preparation of PTFE fibers requires the use of sizing agents which must be removed by vigorous (and proprietary) cleaning processes. PTFE has a notable shortcoming—its susceptibility to degradation by ionizing radiation as experienced during gamma sterilization.32,33 It is completely stable to other forms of sterilization, however. Another potential issue with PTFE concerns the adherence of other materials to its surface. The polymer is noted for its release properties, and this is an advantage for declotting vascular access prostheses.27 However, if a hybrid bio-artificial device requires the bonding of a substantial amount of tissue to the PTFE surface (i.e., without complete mechanical interlock), a potential for delamination may exist. This factor has been recognized and promising approaches to enhance the binding of, for example, endothelial cells are being developed. These include denucleation of ePTFE by saturating its pores with water, use of surfactants and treatment with adhesion molecules (e.g., fibronectin) or adhesion peptides.52,55

473 CH3

Polypropylene(PP)

CH

CH2 n

PP has considerable appeal as a biostable scaffold. It is a relatively strong (tensile strength = 400 MPa), crystalline, high modulus (tensile modulus = 2.6 GPa) thermoplastic in its isotactic form.56 Its hydrocarbon structure renders PP insensitive to hydrolytic attack. However, since every other atom on the polymeric chain is a tertiary carbon, it is susceptible to oxidation during processing, storage and implantation.57 This type of degradation can be effectively countered by the use of antioxidants, which are universally used in PP products.57,58 Consequently, stabilized PP is considered a “permanent” implant material. Biostability, strength, and durability combined with a relatively low inflammatory tendency59-61 make stabilized PP the material of choice for vascular anastomotic sutures,59-61 certain ligament augmentation devices58,62 and mesh for surgical repair.63 PP yarns have been woven into multifilament tubes for investigation as single component vascular prostheses59 or as the reinforcing matrix for partially absorbable composite devices.64 Chronic animal implant studies indicated that PP offered potential advantages in efficacy over expanded PTFE and PET in small diameter vascular applications.59 Biomechanical behavior, which could be readily modulated by varying fiber diameter and weaving conditions, was shown to have a significant effect on tissue ingrowth and the resulting hybrid bio-artificial composite.64,65 PP is one of the least resistant structural polymers to sterilization by ionizing radiation.32,33 However, it is safely sterilized by chemical agents and autoclaving.32-34 Recent innovations in the technology of PP include a syndiotactic product and an ionomer (ion-containing polyethylene)-modified isotactic PP which produce lower moduli (0.5 GPa) than isotactic PP (2.6 GPa).56 Finally, a formulation of polypropylene rendered radiation resistant by the incorporation of hindered amine light stabilizers and a plastomeric ethylene polymer has been described.66 PP is normally considered a low surface energy, hydrophobic material to which surface bonding is difficult. However, certain highly hydrophobic cellular strains adhere strongly to PP.67 This suggests that the adherence of tissue generated on PP scaffolds should be studied on a case by case basis with consideration of the possibility surface modification to control delamination. Polypropylene should be considered in the top tier of candidates for biostable vascular matrices. O

Polyurethanes(PUR)

NHCO

PURs comprise a large family of polymers which is notable for its diversity. The only required attribute that individual members have in common is the presence of the urethane [-NH(CO)O-] group in some repeating sequence on the main chain, from the reaction of an isocyanate group with an alcohol group. Most likely, however, there are other functional groups which make up the soft segment of the PUR, which is generally a copolymer consisting of hard and soft segments. The hard segment generally consists of the

474

Tissue Engineering of Prosthetic Vascular Grafts

Soft Segment Macroglycols

Diisocyanates OCN

CH 2

NCO (MDI)

OCN

CH 2

NCO (HMDI)

OCN

CH 2 NCO (HDI)

OCN

D NCO (DDI)

HO

CH 2 O H (PTMEG) 4 n

HO

CH CH 2 O n H (PPG) CH 3

6

HO Short-chain Diols/Diamines HO

CH 2

4

HO

CH 2

2

HOCH 2

OH (BDO) OH (EG)

O R OCR' O

n

H (polyester diol)

HO

O ROCO H (polycarbonate diol) n

HO

D OH (Dimerol)

CH 2OH (CHDM)

*H2NCH 2CH2NH2 (EDA) *

NH2 (CHDA) NH2 O

*Amines form urea NHCNH linkages by reaction with isocyanates.

Where: R, R' = aliphatic hydrocarbon D = C36 Hydrocarbon from dimerized fatty acid

n = repeating units to produce MWs of ~500-3000

Fig. 43.3 Monomers for polyurethanes.

reaction product of a diisocyanate and a diol or diamine. The soft segment is derived from a macromonomer ranging from several hundred to several thousand in molecular weight.10 Typical monomers for polyurethanes which have been used as “permanent” implants are listed in Figure 43.3.10 Most of the polyurethanes used in medical devices are based on difunctional monomers. Commercial PURs usually contain additives such as catalyst residues, processing aids and stabilizers. These additives often migrate to the surface of the PUR part and serve as the material in contact with blood or soft tissue.10 Certain PURs have demonstrated relatively low thrombogenicity.10 PURs range in physical properties from soft, tough elastomers to strong, rigid structural polymers (e.g., tensile strength = 20-50 MPa, tensile modulus = 5-1150 MPa).10 As a result, PURs have been considered for soft, flexible implantable device components such as pacemaker lead insulation15 tubing and compliant vascular prostheses68-70 as well as hard, rigid components such as pacemaker lead connectors.10 The polyurethanes used in medical devices are usually formed into shapes by solution or melt processes10 and can be solution or melt spun into fibers, cast into porous or solid-wall structures, or extruded into solid-wall tubing. The record for PURs in “permanent” implants is mixed, because of hydrolytic and oxidative degradative mechanisms that may come into play. Generally, the site of

biodegradation is the soft segment (ester, ether, carbonate). While hydrolysis of the urethane (or urea) hard segment linkages is possible, those groups are relatively stable and do not comprise the primary mode of biodegradation. Polyester soft segments have generally degraded quickly and severely by hydrolysis.16 Polether urethanes are susceptible to oxidative degradation as described in the “Modes of Polymer Degradation” section, above.71,72 Since oxidative susceptibility is generally proportional to ether content,9,10,15,16 relatively hard (low ether content) PURs have performed with minimal degradation for periods as long as the lifetime of pacemakers (8-10 yr). Even soft polether urethanes are capable of performing as intended for years in the absence of high stress or strong oxidative attack.15,16 Over the past few years, several industrial concerns have reported on the development of polyurethanes based on polycarbonate soft segments.69,73-75 These have shown very high resistance to degradation in short term studies, although minor hydrolytic effects have been noted on the implant surfaces. 69,76 However, in vascular implants approaching one year, structural degradation of the poly(carbonate urethane) fibers was detectable.69 This is predictive, in my opinion, of a progressive hydrolytic degradation that should make a device designer very wary of choosing poly(carbonate urethanes) as “permanent”

Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses

scaffolds until long term (> 5 years) studies have confirmed their stability. An approach to polyurethane design that uses soft segments composed entirely of hydrocarbon molecules (i.e., a 36 carbon structure derived from dimerized fatty acids) with no ether, ester or carbonate groups to degrade77,78 has theoretical appeal. However, this concept must also pass the test of long term implant stability before it may be used with confidence. Polyurethanes are generally resistant to sterilization of all types except steam and, possibly, dry heat.10 They are susceptible to calcification, especially in highly dynamic applications such as leaflet heart valves.25 Although polyurethanes display some of the most favorable mechanical properties for compliant grafts, no composition has yet demonstrated dependable long term stability in vivo, especially as a fibrous vascular implant. Polyethylene(PE)

CH2 CH2

n

PE is produced in several forms including very low density (VLDPE), low density (LDPE), medium density (MDPE), high density (HDPE), linear low density (LLDPE) and ultra high molecular weight (UHMWPE), each with different thermal and mechanical characteristics.49 The variations in density and physical properties result from differences in polymerization conditions or the use of other olefinic hydrocarbon comonomers.49 All of these molecular structures are amenable to fabrication into fibers or cellular structures, which are the forms most likely to meet the requirements of the “design concept.” Although PE has been reported as being used in vascular prostheses40 and is regularly used in several forms of permanent implants,79 recent literature describing its use in vascular prostheses is sparse. More commonly reported is the use of solid-wall polyethylene tubing with or without coatings for thrombogenicity studies.80-82 The studies show that PE can be modified to improve surface wetability for binding of coatings (e.g., heparin)81 and, potentially, other components for tissue engineered prostheses. Structurally, PE compositions range from relatively low to medium modulus when thermally processed (Table 43.1).83 PE in its various structural forms is stable to hydrolytic media and resistant to oxidation in proportion to its ratio of linear to branched chain structure and crystallinity. PE is susceptible to the phenomenon of environmental stress cracking (ESC) produced by exposure of stressed specimens to aggressive media such as detergents (and, possibly, blood). Cracking occurs because of stress relaxation of “tie mol-

Table 43.1. Properties of PE compositions83 2

Tensile Strength (MPa)

Tensile modulus (MPa)

LDPE MDPE HDPE

15-80 10-20 15-35

55-170 170-380 415-1035

475

ecules” in the zones between crystalline regions in these semicrystalline polymers.84 ESC is reduced in PE with lower crystallinity (LLDPE) or higher molecular weight (UHMWPE) because of higher concentrations of tie molecules relative to crystalline regions.84 Therefore, two forms of PE, LLDPE85 and UHMWPE, have theoretical appeal for ESC resistant fiber-based prostheses. The latter has been used as fibers for sutures.86 The LLDPE is reported to be resistant to cracking.85 A previously unmentioned form of PE deserves mention—gel-spun UHMWPE. These fibers achieve levels of tensile strength (2-4 GPa) and tensile modulus (125-175 GPa) that are multiples of the PE levels reported above.87 Their high stability and strength at very low fiber diameters are potentially useful design features for scaffolds. PE is resistant to chemical sterilization and susceptible to degradation by ionizing radiation.22 Most forms of thermal sterilization would warp the relatively low melting PE (MP 104-135°C).83 The PE family, in my opinion, has not received adequate consideration for use in filament-based vascular prostheses. CH3

O Si

Polydimethylsiloxane (PDMS)

n

CH3

The PDMSs comprise a family of thermoset elastomers with a long history of successful use in implantable devices. 79,88 Structural devices are usually produced by crosslinking 2-part systems at room temperature or higher. The crosslinked products are highly elastic and extensible. The tensile modulus (2-9 MPa) and ultimate tensile strength (2-10 MPa) are relatively low compared to other polymers considered in this chapter.10 However, excellent flex fatigue resistance and tear resistance of some compositions,10 combined with high biostability,88,89 make PDMS a worthy contender for compliant vascular prosthesis scaffolds. PDMS has been studied in vascular prostheses. Because of its low modulus, porous-wall structures have been favored over filamentous configurations. The “Replamineform” process uses sea urchin spine machined into tubular shape as template for fabrication of porous PDMS vascular devices. 90,91 The spine, consisting of microporous calcite, is dissolved in acid, leaving behind an open-cell PDMS structure with adequate strength and the potential for tissue ingrowth.90 PDMS derives much of its strength from silica filler. The filler is actively thrombogenic; however, PDMS without filler can be used as a less thrombogenic coating.82 PDMS has been used as a surface for endothelial cell growth;82 however, it is a relatively low energy surface and the authors caution that resistance to cell detachment under conditions of blood flow must be verified before a hybrid device based on PDMS is deemed suitable for long term function.82 PDMS can be sterilized under relatively mild ionizing radiation conditions and by heat or chemical sterilants.32 A final note on PDMS is in order. The PDMS class of elastomers has come under severe scrutiny in recent years because of its use in mammary prostheses. Reports of autoimmune responses to implanted devices have focused

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on PDMS oil-filled PDMS shell designs. Leakage of the oil is alleged to cause systemic responses to PDMS.92 However, decades of successful use of PDMS in solid form79,93 provide a massive body of evidence that these materials are chemically stable and safe as implants. The list of suppliers of PDMS for “permanent” implants is currently short, but PDMS raw materials are available93 and these elastomers merit serious consideration as scaffolds for “permanent” hybrids. Other Polymer Scaffold Materials This subsection speculates briefly on polymers that have not received much reported consideration for vascular prostheses, but are judged to have the potential to meet the design criteria for a tissue engineered device. The selection is based on my experience in working with or considering the materials for use in “permanent” implants. The sparseness of available information offers the opportunity for technical advancement associated with several risks. At best, there will be substantial development expense to assure successful implementation. There will likely be increased regulatory complexity for new vascular materials. A new vascular material is riskier, because an established clinical history would have revealed characteristics of a material that might not be anticipated in the development plan. Finally, there is the risk that some of these materials have been studied by others and rejected for good cause without the results being published. The subsection is now separated into two parts, the first listing high-modulus linear polymers which are best utilized in fiber form, the second listing low-modulus elastomers which may be fabricated into fiber-based, solid-wall or porous devices. High Modulus Polymers PTFE is just one of several polymers which belong to the fluoropolymer category. It also includes: poly(ethylene tetrafluoroethylene), poly(ethylene chlorotrifluoroethylene), poly(vinylidine fluoride), fluorinated ethylene propylene, perfluoroalkoxy resin, polychlorotrifluoroethylene and others.49,94 They all offer exceptional resistance to hydrolysis and oxidation. Just as with PTFE, they are sterilizable by all common means except, possibly, ionizing radiation.32 The “other” fluoropolymers are worth considering as an alternative to PTFE because most are more readily proccessable by melt or solution techniques. They also offer a broad range of modulus characteristics to choose from.94 Several rigid polyamides should have the chemical resistance to provide long term service in vivo. These include Nylon 11, Nylon 1295 and aromatic polyamides (Aramids).96,97 Their stability arises either from hydrophobicity of the backbone (Nylon 11, 12) or high crystallinity (Aramid). Braided Aramid fiber prostheses have been studied for artificial ligaments.96,97 While mechanical degradation was observed under dynamic conditions,96 long term chemical stability in human plasma was confirmed. Certain aromatic polyesters and poly(ester amides) belong to a class of polymers called liquid crystalline poly-

Tissue Engineering of Prosthetic Vascular Grafts

mers (LCPs). They are so named because of their high degree of order in the liquid phase. Upon cooling, these polymers solidify to structures with anisotropic skin-core morphologies and tensile properties among the highest of known polymers (i.e., tensile strengths to 240 MPa, tensile moduli to 32 GPa).98 Chemical degradation resistance is also exceptionally high. The materials can be readily extruded into tubing and fibers.99 If the fibers are not too stiff, LCPs may be worth considering as structural matrices for hybrid vascular prostheses. A group of rigid thermoplastics having amorphous or semicrystalline morphologies and superior resistance to chemical degradation includes the polysulfones (polysulfone, polyethersulfone, polyarylsulfone) the polyketones (PEK, PEEK, PAEK, PEKK, where P=poly; A=aryl; E=ether; K=ketone), poly(ether imide), polyimide, poly(methyl methacrylate) and others.98,100-104 The stated polymers normally would comprise the nonfibrous component of a fiber-reinforced composite used, for example, in orthopedic devices.100-104 Investigation of fibrous forms of these materials is a worthwhile undertaking, in my opinion. This is being done with poly(methyl methacrylate) currently, and the results are impressive.103 For example, the ultimate tensile strength of PMMA can be increased from 25-50 MPa to as high as 220 MPa while enhancing ultimate elongation from 5-35% with fiber drawdown ratios of ~20.103,104 Tensile moduli increase from ~2 GPa in castings to ~8 GPa in the fibers, which would suggest the use of very thin fiber diameters in vascular scaffolds. Poly(acrylonitrile-co-vinyl chloride) has been fabricated into hollow fiber form as a cell encapsulation membrane for hybrid bio-artificial organs and hemo-filtration membranes. It has been shown to be stable in long term implants and can likely be extruded or solution spun into fibers suitable for consideration for vascular scaffolds.105 Low Modulus Elastomers Elastomers meriting consideration in this subsection, in addition to the ones already considered, may be either thermoplastic or thermoset. Both types may be fabricated into fibrous, porous or solid wall form. Thermoset rubbers are often prepared from prepolymers (“millable gums”) by hot molding processes. In such cases, residues from the “vulcanization” (crosslinking) process may need to be removed by extraction to make the scaffold nontoxic. The elastomers are often compounded with fillers, stabilizers, processing aids and other ingredients to control physical and chemical properties of the final product. Certain of the compositions to be described may not be available in prepolymer form for fabrication into desired shapes, but are sold in finished form by the materials manufacturer. Therefore, any evaluation of a scaffold configuration would require the approval and fabrication by the material supplier. In these times of high liability risk, some elastomer suppliers may not approve the use of their product as “permanent” implants. Nonetheless, elastomers offer the potential advantage of radial compliance and deserve consideration. This sec-

Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses

tion provides a very brief overview, but recent reviews of elastomers are more detailed and comprehensive.106,107 Polyolefins include rubbers synthesized from olefinic hydrocarbons, generally crosslinked with dienes. Rubbers of ethylene, propylene, isobutylene and higher olefins are included here. They may be compounded to various ranges of strength, hardness and elongation. They are highly resistant to hydrolysis, and have low residual olefin content for oxidation stability (in contrast to rubbers from dienes such as butadiene). However, they have branched-chain sites which may be susceptible to irradiation and thermooxidative degradation, similar to polypropylene. One polyolefin, poly(1-hexene) crosslinked with methyl hexadiene, has enjoyed special success in implantable devices (finger joints, artificial heart pumps, compliance chambers, intervetrebral discs).106 It has excellent biocompatibility (after solvent extraction), fatigue resistance and biostability. Fluoroelastomers and fluorosilicone elastomers should have excellent resistance to processing, storage and biological media while retaining low modulus and resilience of true elastomers.106,107 Polychloroprene, a diene rubber, will not hydrolyze and has much greater resistance to oxidation than other diene rubbers (e.g., natural rubber, butadiene rubber). It should be quite biocompatible if extracted after vulcanization. Many other elastomers are available commercially which may be worth considering as scaffolds.106,107 The ones selected for discussion are the ones I might select in a first round of testing. All of those listed in this subsection would require substantial process development and validation before clinical studies would be justified.

Concluding Comments The extensive list of polymers suggested as potential scaffolds for hybrid bio-artificial vascular prostheses was feasible because the “design concept” was rather broad and limited in detail. With a more stringent set of design criteria in place, some of the polymer candidates should be eliminated. The strategy of this chapter was to be inclusive, so that the major categories of relatively stable polymers were captured. One category and one polymer within that category will be the best one. The search for that biomaterial offers certain challenge but great potential reward. Finally, I offer a biased opinion about the hybrid device discussed in this chapter. Ideally, the permanent scaffold is an interim solution on the road to a resorbable scaffold which would produce a completely biological, fully functional neo-artery. My personal experience with a degradable cellular polyurethane vascular prosthesis was that a biological conduit indeed was generated. It had many of the structural elements of natural artery, but they were not organized in the same way. The conduit dilated to unacceptable levels over several months. Dilatation was also observed by other investigators using this type of material.108 A successful result requires guided tissue regeneration which withstands the pressures of the arterial system for the long term. The answer will likely lie in the structural design of the temporary scaffold, its mechanical properties and its degradation characteristics. The pursuit of this design is,

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indeed, a worthy one which I believe will ultimately be successful. Acknowledgments The literature provided by Drs. Giovani Galletti, Robert Guidoin, Howard Greisler, Vince Medenhall, Colin Pitt and J. Paul Santerre is gratefully acknowledged. I sincerely thank Mrs. Sandra Brigham for her expert preparation of this manuscript. References 1. Stoeger KJ. Implementing the new quality system requirements. Design controls. Biomed Instrum Technol 1997; 31(2):119-127. 2. AAMI Standards and Recommended Practices. Biological evaluation of medical devices. Washington, D.C.: Assoc for the Advancement of Med Instrumentation, 1996:4. 3. Graham L, Whittlesey D, Bevacqua B. Cardiovascular implantation: Vascular grafts. In: Ratner B et al, eds. Biomaterials Science, An Introduction to Materials in Medicine. San Diego: Academic Press, 1996:420-422. 4. Kantor C. Biological small diameter vascular grafts. AAMI Medical Device Research Report 1996; 3(1):8-11. 5. Galletti PM. Biomaterials availability in the U.S. J Biomed Mater Res 1996; 32(3):289-291. 6. Citron P. Medical devices: Factors adversely affecting innovation. J Biomed Mater Res 1996; 32(1):1-2. 7. O’Connor KW. Biomaterial bill passes Congress, signed into law. The AIMBE News, Fall, 1998; 6(3):1,2. 8. Coury AJ. Preparation of specimens for blood compatibility testing. Cardiovasc Pathol 1993; 2(3) (Suppl):1015-1105. 9. Coury AJ. Chemical and biochemical degradation of polymers. In: Ratner B et al, eds. Biomaterials Science, An Introduction to Materials in Medicine. San Diego: Academic Press, 1996; 243-260. 10. Coury AJ, Slaikeu P, Cahalan P et al. Factors and interactions affecting the performance of polyurethane elastomers in medical devices. J Biomater Appl 1988; 3:130-179. 11. Zhu KJ, Hendren RW, Jensen K et al. Synthesis, properties and biodegradation of poly(1,3-trimethylene carbonate). Macromolecules (1991); 24:1736-1740. 12. Santerre JP, Labow RS, Adams GA: Enzyme-biomaterial interactions: Effect of biosystems on degradation of polyurethanes. J Biomed Mater Res 1993; 27:97-109. 13. Ratner BD, Gladhill KW, Horbett TA. In vitro studies of the enzymatic biodegradation of polether urethanes. Trans 12th Ann Mtg Soc Biomater 1986; 9:190. 14. Ratner BA, Tyler BJ. Variations between biomer lots 2: The effect of differences between lots on in vitro enzymatic and oxidative degradation of a commercial polyurethane. J Biomed Mater Res 1993; 27:327-334. 15. Stokes K, Coury A, Urbanski P. Autooxidative degradation of implanted polyether polyurethane devices. J Biomater Appl 1987; 1:412-448. 16. Coury AJ, Stokes KB, Cahalan PT et al. Biostability considerations for implantable polyurethanes. Life Support Systems 1987; 5:25-39. 17. Sutherland K, Mahoney JR, Coury AJ et al. Degradation of materials by phagocyte-derived oxidants. J Clin Invest 1993; 92:2360-2367. 18. Santerre JP, Labow RS, Duguay DG et al. Biodegradation evaluation of polyether- and polyester-urethanes with

478 oxidative and hydrolytic enzymes. J Biomed Mater Res 1994; 28:1187-1199. 19. Taylor G, Gsell R, King R et al. Stability of N2 packaged gamma irradiated UHMWPE. Trans 23rd Ann Mtg Soc Biomater 1997; 20:421. 20. Furman BD, Reish TG, Li S. The effect of implantation on the oxidation of ultra high molecular weight polyethylene. Trans 23rd Ann Mtg Soc Biomater 1997; 20:427. 21. Wang A, Polineni VK, Essner A et al. Effect of radiation dosage on the wear of stabilized UHMWPE evaluated by hip and knee joint simulators. Trans 23rd Ann Mtg Soc Biomater 1997; 20:394. 22. Premnath V, Harris WH, Jasty M et al. Gamma sterilization of UHMWPE articular implants: An analysis of the oxidation problem. Biomaterials 1996; 17:1741-1753. 23. del Prever EB, Crova M, Costa L et al. Unacceptable biodegradation of polyethylene in vivo. Biomaterials 1996; 17:873-878. 24. Lewis G. Polyethylene wear in total hip and knee arthroplasties. J Biomed Mater Res 1997; 38(1):55-75. 25. Pathak Y, Schoen FJ, Levy RJ. Pathologic calcification of biomaterials. In: Ratner B et al, eds. Biomaterials Science, An Introduction to Materials in Medicine. San Diego: Academic Press, 1996; 272-281. 26. Joshi RR, Frautschi JR, Phillips RE. Immobilized heparin and heparin-bisphosphonate prevent polyurethane calcification and thrombosis: In vitro and in vivo studies. 5th World Biomater Congress 1996; May 29-June 2: 610. 27. Ku DN, Allen RC. Vascular grafts. In: Bronzino JD, ed. The Biomedical Engineering Handbook. Boca Raton: CRC Press, Inc 1995; 1871-1878. 28. Lawton EL, Ringwald EL. Physical constants of Poly(oxyethylene oxyterephthaloyl)[poly(ethylene terephthalate)]. In: Brandrup J., Immergut EH, eds, Polymer Handbook. 3rd Ed, New York: John Wiley & Sons, 1989:V101-V105. 29. Williams DF. Review: Biodegradation of surgical polymers. J Mater Sci 1982; 17:1239-1240. 30. Kopecek J, Ulbrich K. Biodegradation of biomedical polymers. Prog Polym Sci 1983; 9:31. 31. Ambrosio L, Apicella A, Mensitieri M et al. Physical and chemical decay of prosthetic ACL after in vivo implantation. Clin Mater 1994; 15:29-36. 32. Lee HB, Kim SS, Khang G. Polymeric biomaterials: Sterilization. In: Bronzino JD, ed. The Biomedical Engineering Handbook. Boca Raton: CRC Press, Inc 1995; 581-592. 33. Kowalski JB, Morrissey RF. Sterilization of implants. In: Ratner, B et al, eds. Biomaterials Science, An Introduction to Materials in Medicine. San Diego: Academic Press 1996; 415-420. 34. Booth AE. Industrial sterilization technologies: New and old trends shape manufacturer choices. Med Device Diagn Industry; 1995; Feb: 64-72. Author: Please provide volume number rather than month. 35. Zenni GC, Ellinger J, Lam T et al. Biomaterial-induced macrophage activation and monokine release. J Invest Surg 1994; 7:135-141. 36. Greisler HP, Petsikas D, Cziperle DJ et al. Dacron stimulation of macrophage transformtin growth factor-β release. Cardiovasc Surg 1996; 4(2):169-173. 37. Greisler HP, Dennis JW, Endean ED et al. Derivation of neointima in vascular grafts. Circ 1988; 78(Suppl 1):I-6 to I-12. 38. Swartbol P, Truedsson L, Parsson H et al. Tumor necrosis factor-α and interleukin-6 release from white blood

Tissue Engineering of Prosthetic Vascular Grafts cells induced by different graft materials in vitro are affected by pentoxifylline and iloprost. J Biomed Mater Res 1977; 36: 400-406. 39. Greisler HP, Schwarcz TH, Ellinger J et al. Dacron inhibition of arterial regenerative activities. J Vasc Surg 1986; 3(5):747-756. 40. Peng T, Gibula P, Yao K et al. Role of polymers in improving the results of stenting in coronary arteries. Biomaterials 1996; 17(7):685-694. 41. Murphy JG, Schwartz RS, Edwards WD et al. Percutaneous polymeric stents in porcine coronary arteries. Circ 1992; 86:1596-1604. 42. Marois Y, Guidoin R, Roy R et al. Selecting valid in vitro biocompatibility tests that predict in vivo healing response of synthetic vascular prostheses. Biomaterials 1996; 17:1835-1842. 43. Marois Y, Chakfe N, Guidoin R et al. An albumin-coated polyester arterial graft: In vivo assessment of biocompatibility and healing characteristics. Biomaterials 1996; 17:3-14. 44. Cziperle DJ, Joyce KA, Tattersall CW et al. Albumin impregnated vascular grafts: Albumin resorption and tissue reactions. J Cardiovasc Surg 1992; 33:407-414. 45. White JE, Werkmeister JA, Edwards GA et al. Structural analysis of a collagen-polyester composite vascular prosthesis. Clin Mater 1993; 14:271-276. 46. Fournier N, Doillon CJ. Biological molecule-impregnated polyester: An in vivo angiogenesis study. Biomaterials 1996; 17:1659-1665. 47. Greisler HP, Klosak I, Dennis JW et al. Endothelial cell growth factor attachment to biomaterials. Trans Am Soc Artif Intern Organs 1986; XXXII:346-349. 48. Phaneuf MD, Quist WC, Bide MJ et al. Modification of polyethylene terephthalate (Dacron) via denier reduction: Effects on material tensile strength, weight, and protein binding capabilities. J Appl Biomater 1995; 6:289-299. 49. Lee HB, Kim SS, Khang G. Polymers used as biomaterials. In: Bronzino JD, ed. The Biomedical Engineering Handbook. Boca Raton: CRC Press, Inc 1995; 580-591. 50. Shalaby S. Classes of materials used in medicine: Fabrics. In: Ratner B et al, eds. Biomaterials Science, An Introduction to Materials in Medicine. San Diego: Academic Press 1996; 118-124. 51. Greisler HP. Growth factor release from vascular grafts. J Controlled Release 1996; 39:267-280. 52. Greisler HP, Johnson S, Joyce K et al. The effects of shear stress on endothelial cell retention and function on expanded polytetrafluoroethylene. Arch Surg 1990; 125:1622-1625. 53. Changdran KB. Blood-interfacing implants. In: Bronzino JD, ed. The Biomedical Engineering Handbook. Boca Raton; CRC Press, Inc 1995; 648-655. 54. Yoganathan AP. A brief history of heart valve prostheses. In: Bronzino JD, ed. The Biomedical Engineering Handbook. Boca Raton: CRC Press, Inc 1995; 1848-1853. 55. Wigod MD, Klitzman B. Quantification of in vitro endothelial cell adhesion to vascular graft material. J Biomed Mater Res 1993; 27:1057-1062. 56. Liu CK. Medical fibers spun from polypropylene. Proc 13th Southern Biomed Eng Conf 1994; 748-751. 57. Williams DF. Review: Biodegradation of surgical polymers. J Mater Sci 1982; 17:1233-1237. 58. Gibbons DF, Mendenhall HV, Van Kampen CL et al. The effect of motion on the tissue response to polymeric fiber implants. Clin Mater 1994; 15:37-41.

Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses 59. Greisler HP, Tattersall CW, Henderson SC et al. Polypropylene small-diameter vascular grafts. J Biomed Mater Res 1992; 26:1383-1394. 60. Bakkum EA, Dalmeijer RAJ, Verdel MJC et al. Quantitative analysis of the inflammatory reaction surrounding sutures commonly used in operative procedures and the relation of postsurgical adhesion formation. Biomaterials 1995; 16:1283-1289. 61. Faulkner BC, Tribble CG, Thacker JG et al. Knot performance of polypropylene sutures. J Biomed Mater Res (Appl Biomater ) 1996; 33:187-192. 62. McPherson GK, Mendenhall HV, Gibbons DF et al. Experimental mechanical and histologic evaluation of the Kennedy ligament augmentation device. Clin Orthop 1985; 196:186-195. 63. Bellon JM, Contreras LA, Bujan J et al. Effect of phosphatidyl choline on the process of peritoneal adhesion following implantation of a polypropylene mesh prosthesis. Biomaterials 1996; 17:1369-1372. 64. Greisler HP. Effects of polypropylene’s mechanical properties on histological and functional reactions to polyglactin 910/polypropylene vascular prostheses. Am Coll Surg Surg Forum 1987; XXXVIII:323-326. 65. Zenni GC, Gray JL, Appelgren EO et al. Modulation of myofibroblast proliferation by vascular prosthesis biomechanics. ASAIO J (1993); Vol. 39, No. 3:M496-M500. 66. Portnoy R. Clear, radiation-tolerant, autoclavable polypropylene. Med Plast Biomater 1997; 4(1):40-48. 67. Kiremitci-Gumusderelioglu M, Pesmen A. Microbial adhesion to ionogenic PHEMA, PU and PP implants. Biomaterials 1996; 17:443-449. 68. DeCossart L, Annis D. An assessment of rigidly controlled technique for the implantation in dogs of a new microfibrous 3.8 mm I.D. polyurethane arterial prosthesis. Proc 2nd Ann Scientific Session, Acad Surg Res 1986; Oct. 31-Nov. 1, Clemson SC: 36. 69. Zhang Z, Marois Y, Guidoin RG et al. Vascugraft® polyurethane arterial prosthesis as femoro-popliteal and femoro-peroneal bypasses in humans: Pathological, structural and chemical analyses of four excised grafts. Biomaterials 1997; 18:113-124. 70. Doi K, Matsuda T. Enhanced vascularization in a microporous polyurethane graft impregnated with basic fibroblast growth factor and heparin. J Biomed Mater Res 1997; 34:361-370. 71. Zhao QH, McNally AK, Rubin KR et al. Human plasma α2-macroglobulin promotes in vitro oxidative stress cracking of Pellethane 2363-80A: In vivo and in vitro correlations. J Biomed Mater Res 1993; 27:379-389. 72. Schubert MA, Wiggins MJ, Anderson JA. Role of oxygen in biodegradation of poly(etherurethane urea) elastomers. J Biomed Mater Res 1997; 34:519-530. 73. Szycher M, Edwards A, Carson RJ. In vivo testing of a biodurable polyurethane. Trans 23rd Ann Mtg Soc Biomater 1997; 296. 74. Ward RS, White KA, Gill RS. Development of biostable thermoplastic polyurethanes with oligomeric polydimethylsiloxane end groups. Trans 21st Ann Mtg Soc Biomater 1995; 18:268. 75. Kato YP, Dereume JP, Kontges H et al. Preliminary mechanical evaluation of a novel endoluminal graft. Trans 21st Ann Mtg Soc Biomater 1995; 18:81. 76. Mathur AB, Collier TO, Kao W et al. In vivo biocompatibility and biostability of modified polyurethanes. J Biomed Mater Res 1997; 36:246-257.

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77. Coury AJ, Hobot CM, Iverson VB. Novel soft segment approaches to implantable biostable polyurethanes. Trans 4th World Biomater Cong 1992; 661. 79. Visser SA, Hergenrother RA, Cooper SL. Classes of materials used in medicine: Polymers. In: Ratner B et al, eds. Biomaterials Science, An Introduction of Materials in Medicine. San Diego: Academic Press 1996; 50-60. 80. Rubens FD, Weitz JI, Brash JL et al. The effect of antithrombin III-independent thrombin inhibitors and heparin on fibrin accretion onto fibrin-coated polyethylene. Thromb Hemost 1993; 69(2):130-134. 81. Evangelista RA, Sefton MV. Coating of two polyetherpolyurethanes and polyethylene with a heparin-poly(vinyl alcohol) hydrogel. Biomaterials 1986; 7:206-211. 82. Keough EM, Mackey WC, Connolly R et al. The interaction of blood components with PDMS [polydimethylsiloxane and LDPE (low-density polyethylene)] in a baboon ex vivo arteriovenous shunt model. J Biomed Mater Res 1985; 19:577-587. 83. Quirk RP, Alsamarraie MAA. Physical constants of poly(ethylene). In: Brandrup J and Immergut EH, eds. Polymer Handbook. 3rd Ed. New York: John Wiley & Sons, 1989:V15-V26. 84. Lustiger A. Understanding environmental stress cracking in polyethylene. Med Plast Biomaterials 1996; 3(4):12-18. 85. Thermoplastic Processes, Inc. LLDPE tubing resists cracking. Med Plast Biomater 1996; 3(3):61. 86. Tomita N, Tamai S, Morihara T et al. Handling characteristics of braided structure materials for tight tying. J Appl Biomater 1993; 4:61-65. 87. Prevorsek DC. Preparation, structure, properties and applications of gel-spun ultrastrong polyethylene fibers. Trends Pol Sci 1995; 3(1):4-11. 88. McMillin CR. An assessment of elastomers for biomedical applications. In: Szycher, M, ed. High Performance Biomaterials. Lancaster: Technomic Publishing Co, Inc 1991; 37-49. 89. Chawala AS, Hinber I. Laboratory evaluation of explanted gel filled silicone breast implants. Trans World Biomater Cong 1996:300. 90. Tizian C, Salyer KE. Production process of a microvascular prosthesis using the “replamineform” principle. Int J Artif Organs 1980; 6:364-465. 91. Hiratzka LF, Goeken JA, White RA et al. In vivo comparison of replamineform silastic and bioelectric polyurethane arterial grafts. Arch Surg 1979; 114(6):698-702. 92. Nicholson J III, Hill SL, Frondoza CG et al. Silicone gel and octamethylcyclotetrasiloxane [D4] enhances antibody production to bovine serum albumin in mice. Trans 5th World Biomater Cong 1996; 304. 93. Winn A. Factors in selecting medical silicones. Med Plast Biomater 1996; 3(2):16-19. 94. Resins and compounds: Fluoroplastics. In: Kaplan WA, managing ed. Modern Plast Encyclopedia ’97. New York: McGraw-Hill Co’s., Inc., 1996:B158-B160. 95. Resins and compounds: polyamides. In: Kaplan WA, ed. Modern Plast Encyclopedia, ‘97. New York: McGraw-Hill Co., Inc., 1996:B175. 96. Durselen L, Claes L, Ignatius A et al. Comparative animal study of three ligament prostheses for the replacement of the anterior cruciate and medical collateral ligament. Biomaterials 1996; 17(10):977-982. 97. Wening JV, Lorke DE. A scanning electron microscope (SEM) investigation of Aramid (Kevlar) fibers after incubation in plasma. Clin Mater 1992; 9:1-5.

480 98. Canale B, Hanley S, Braeckel M. New possibiliities for liquid crystal polymers. Med Plast Biomater 1995; 2(3):24-31. 99. ACT Medical, Inc. Liquid crystal polymer tubing. Med Device Diagn Industry 1997; 19:211. 100. Moore R, Beredjiklian P Rhoad R et al. A comparison of the inflammatory potential of particulated derived from two composite materials. J Biomed Mater Res 1997; 34:137-147. 101. Zhang G, Latour Jr RA, Kennedy JM et al. Long-term compressive property durability of fiber reinforced polyetheretherketone composite in physiological saline. Biomaterials 1996; 17(8):781-789. 102. Barton AJ, Sagers RD, Pitt WG. Bacterial adhesion to orthopedic implant polymers. J Biomed Mater Res 1996; 30:403-410. 103. Gilbert JL, Ney DS, Lautenschlager EP. Self-reinforced composite poly(methyl methacrylate): static and fatigue properties. Biometerials 1995; 16:1043-1055.

Tissue Engineering of Prosthetic Vascular Grafts 104. Lewis G. Properties of acrylic bone cement: state of the art review. J Biomed Mater Res (Appl Biomater) 1997; 38:155-182. 105. Shoichet MS, Rein DH. In vivo biostability of a polymeric hollow fiber membrane for cell encapsulation. Biomaterials 1996; 17(3):285-290. 106. McMillin CR. An assessment of elastomers for biomedical applications. In: Szycher M, ed. High Performance Biometerals. Lancaster: Technomic Publishing Co., Inc., 1991:37-49. 107. Courtney PJ, Sevenson JA, Verosky C. Bonding elastomers with adhesives. Med Plast Biomater 1997; 4(3):60-68. 108. Galletti G, Farruggia F, Baccarini E et al. Prevention of platelet aggregation by dietary polyunsaturated fatty acids in the biodegradable polyurethane vascular prosthesis: An experimental model in pigs. Ital J Surg Sci 1989; 19(2):121-130.

Scaffold Engineering Material Aspects

CHAPTER 44 Biophilic Polymers: What’s on the Horizon? Patrick T. Cahalan

T

his chapter was outlined for a section of this book entitled ‘Bio-Interactive’ Prostheses, and was further subdivided to a section including biostable polymers/materials. The other chapter in this subsection, titled “Biostable Polymers as Durable Scaffolds for Tissue Engineered Vascular Prostheses” is being written by a dear colleague, Art Coury, who shares this author’s sense of trepidation as to achieving substance in the context of the remaining chapters of the overall section. It is hoped that, although this chapter will be written from an industrial perspective, it will have some practical value for the reader. The word biophilic cannot be found in the dictionary, but then neither can the word biocompatible. In the early 1980s materials were placed in four basic groups: 1. Bioinert; 2. Biocompatible; 3. Bioactive; and 4. Biointeractive. Bioinert surfaces were hypothesized to be “invisible” to the body, and to have little or no interaction. Proposed early examples were negative surface charge to repel cell adhesion,1 high surface energy materials2 such as pyrolytic carbon and low critical surface tension materials3 such as fluoropolymers, which also claimed low protein and platelet adhesion. Also in the early 1980s, hydrophilic materials were suggested to be low protein adsorbing and platelet adhering,4 and in particular PEG-like surfaces were claimed to have increased surface motion (flagella-like activity) that served to decrease protein adsorption and denaturation.5 The term biocompatible could be applied to any material that when implanted showed an equal or better tissue response compared to a control material, such as polypropylene, that was accepted as biocompatible. Bioactive surfaces were surfaces designed to promote an advantageous effect such as preferential adsorption of albumin by C-18 alkylated surfaces.6 Finally, biointeractive surfaces were those designed to interact with a specific physiological mechanism such as coagulation; an early example is ionically bound heparin for local release to decrease coagulation.7,8 Heparin was finally covalently immobilized and shown to have bioactivity without releasing by Larm;9 this surface has been shown in vitro to bind ATIII to produce a catalytic rate of thrombin deactivation. These categories were arrived at largely from the perspective of the empirical surface response, and arguably need some revaluation with more instrument-intensive analytical methods now available to properly order them by rank in a new era of tissue engineered materials. The editors have chosen new labels, and have done so with the healing response in mind. Thus they have set the stage for the conceptual advantages of tissue engineering. In labeling bio-inert materials as those exhibiting insufficient healing, biolized (endothelialized surfaces) as surface healing, and bio-interactive as complete healing, they have set the goal to incorporate all that has been learned to date towards designing materials and structures Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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that should more closely approximate naturally functioning tissue. In the early years of vascular graft implantation, Voorhees’ observations with Vinyon “N” cloth in dog aortas clearly showed that a synthetic material could serve as a conduit, and be accepted by the tissue.10 Tissue engineering, or the integration of biology and biomaterials, holds out the promise of producing more than an accepted conduit, and obviating all the problems associated with the current commercial products such as thrombotic occlusion, anastomotic hyperplasia, aneurysmal dilatation, and infection. Presumably, bio-interactive prostheses are the desired approach to achieving this goal. During the period between 1982 and 1988, this author was involved with research to find a small diameter (4 mm or less) vascular graft that could exhibit patency at least equivalent to that of autologous saphenous vein. In reality the effort was more a development, or an evaluation exercise, that included animal implantation of numerous technology platforms to include: 1. Biodegradable polyurethane-PLLA grafts; 2. Biostable polyurethane grafts made by phase inversion, electrostatic spinning, and spray techniques; 3. Alternatively (non glutaraldehyde) fixed heterografts; 4. Plasma TFE coated Dacron prostheses; 5. Silicone replamarinaform porous conduits; and 6. Dacron and PTFE control grafts from commercial sources. Predominant modes of failure were: aneurysmal dilatation (1 and 2); in vivo degradation across species (3); rejection by surgeons at the need to cut the prostheses with an electrical device (the flaming graft), and undesirable suture characteristics (4); poor crush resistance leading to thrombotic occlusions, and poor suture pull out properties (5). The Dacron and PTFE grafts had patency rates not competitive with autologous vein, and in general showed a fibrotic nonintegrated histology. Fabricating prostheses for implantation gives one an appreciation for necessary mechanical properties such as suture strength, ease of suture, burst strength, porosity, and wall thickness for matching vessel anastomosis. Since these properties have been achieved for the most part in commercially available materials, it was felt that applying surface modification technologies to existing materials might be able to control problems such as thrombosis and anastomotic hyperplasia, and possibly achieve a healing response that would result in neovascularization, endothelialization, and a tissue response that was not primarily a fibrotic or a chronic inflammatory response. The latter could help a graft towards long term patency and infection resistance. If one defines biophilic polymers as those specifically synthesized to interact in an advantageous manner with the body, there are many claims, but few real candidates are on the near horizon for application. There are numerous claims for heparinoid polymers.21 These polymers often have carboxyl and sulfate functionality approximating the ratio contained in heparin. Studies showing less adhered platelets, or changing clotting times, are generally used as proof of heparin-like activity. In light of studies demonstrating the unique pentasaccharide sequence required for activation of ATIII, and slight changes in this structure resulting in 90% to com-

Tissue Engineering of Prosthetic Vascular Grafts

plete loss of heparin activity, it is highly optimistic to hope for a synthetic anionic polymer with heparin properties.22 Studies in our laboratory with compounds reported as heparinoid or anticoagulant fail to produce heparin-like activity when measured in solution with ATIII and thrombin. Surfaces with immobilized cationic or anionic character can adsorb ATIII and thrombin respectively, and give false positives for heparin activity. An amine functional surface can show binding of ATIII that is resistant to extensive 0.15 ionic strength rinsings, and if given time will deactivate thrombin, though not in a catalytic fashion. An anionic surface can bind thrombin that does not rinse off in the normal rinse steps used before difference measurement using chromogenic substrate, giving a false measure of thrombin deactivated. Active thrombin can be witnessed by adding substrate back to the surface and obtaining a color change in the substrate. Since it is well known that anionic surfaces are contact activating, it is of concern to have an anionic surface present that does not have high heparin activity. Since coagulation is so dependent on platelets, it is hypothesized that by using a polymer that mimics the nonactivated platelet membrane, namely phosphatidylcholine, it would be possible to prevent platelet interaction with the polymer surface (Fig. 44.1). Phosphatidylcholine in the form of commercially available lecithin can coat surfaces of hydrophobic polymers, but is easily removed with flowing plasma. In attempt to remedy this problem, commercial efforts have given rise to a network polymer with strong adhesive properties to polymers, and containing pendant phosphorylcholine moieties as seen in Figure 44.2.23 The most common network polymer is an acrylate backbone that has good adhesive properties for PVC and polyurethanes. The polymer gives excellent wetting properties to the surfaces it coats, and has very low platelet adhesion. Human blood testing in our laboratories show this polymer to be promising, but material-dependent. Perhaps the most innovative biophilic polymer synthesis to date is the work of David Tirrell in recombinant artificial structural proteins. These polymers hold out the potential of engineering proteins with controlled crystallinity and expression of specific peptide sequences at surfaces. If desirable mechanical properties can be engineered into the protein polymers, and if processing methods are devised to make fibers, extrusion, or coating possible, then these polymers can make their way into devices. There has been somewhat of a rebirth in surface modifying additives (SMAs). If a commercial group could blend in additives to polymers that would express themselves on surfaces of devices that their customers manufacture, and demonstrate enhanced biological response, then a premium price could be charged for such materials. To date, because of the lack of tonnage of polymers used in biomaterials, it has been less than attractive for larger chemical companies to specifically manufacture biomaterials. That, together with the liability issues surrounding medical devices, makes it difficult to see new formulations coming from the larger chemical companies. SMA technology is particularly being

Biophilic Polymers: What’s on the Horizon?

483

Fig. 44.1. Membrane lipid—phosphatidylcholine.

Fig. 44.2. Phosphorylcholine surface.

applied for blood-contacting surfaces for short term use, such as cardiopulmonary bypass circuits. For the most part, biophilic polymers on the near horizon are going to be polymers that have been surface modified. Simple adsorbed coatings will be problematic from the standpoint of assuring stability for permanent implants, and there continues to be a quest for the universal biophilic polymer that can coat all materials. It was the choice of our group to attempt to comprehensively evaluate methods to covalently couple molecules to the surfaces of common polymeric biomaterials in hopes of creating new modified biophilic polymers that are stable and have controlled loading of bioactive molecules. If we achieve covalent and stable

attachment, then, with some license in nomenclature, we have created new polymers. In light of later chapters on surface modification for specific surfaces, we hope that it is appropriate in this chapter to discuss the more general methods of surface modification that can be used to get to the biophilic surface, and point to critical concerns and opportunities that may be afforded from our experiences. In 1989 we created a biomaterials laboratory in Maastricht, The Netherlands. The main objective was to focus on surface modification technologies to improve the biocompatibility of materials used to construct biomedical devices. The industrial laboratory was ideally situated 100 meters from the Biomedical Research Institute of The

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University of Maastricht, and the group of Professor Coen Hemker, an experienced and respected expert in thrombosis and hemostasis. A partnership was established between the University of Maastricht (RL), The Bakken Research Center (BRC) and The Center for Surface and Materials Analysis (CSMA), in Manchester, England. The BRC was to provide modified biomaterials for characterization by CSMA, using predominantly XPS and TofSIMS, and biocompatibility testing by RL. The partnership received a three year research subsidy from the EC in the form of a Brite EuRam grant (BE 5972). In this project a broad survey of physical and chemical methods to modify biomaterials was performed (Table 44.1), and these methods were evaluated on seven common biomaterials for feasibility. 11 Equipped in our laboratory with a plasma reactor, corona treater, ozone generator, UV photopolymerization equipment, and a nearby gamma-beta commercial irradiation source, we were prepared to generate many of the surfaces proposed to be more biocompatible for head to head comparisons with our partners’ assistance. What will be presented in this chapter are our somewhat narrow perspectives on making polymers biophilic; our experience with synthesis is limited to modifying the surface of biomaterials, and does not include creating new bulk polymers. With respect to the latter, we maintain surveillance of the literature and sample new materials through our testing schemes when available from the developers. The processes mentioned in Table 44.1, although not complete, represent treatments that were performed on at least 3-7 common materials for each method. Some brief comments will be given here that are felt to be important concerning these surface modification methods and their applicability to biomaterials and biomedical devices.

Physical Methods Irradiation Early work by Ratner and Hoffman4 showed the potential of grafting biomaterials with irradiation sources. Our efforts included irradiation using electron beam sources to successfully graft to polyolefins. E-beam or gamma radiation can introduce several functionalities to polymeric materials, such as free radicals, vinyl bonds and, in the presence of oxygen, hydroperoxides, carbonyls, aldehydes and car-

Table 44.1. Methods to functionalize/activate surfaces (Brite EuRam) Physical

Chemical

Irradiation Plasma Corona CVD and PVD “Simple” Coating Entanglement Textruing Ion Beam Implant

Oxidation Reduction Hydrolysis Ozonization Silanization Grafting Photocuping Add’n/Subst’n

boxyl groups. This can also be accompanied by crosslinking or chain scission, the former leading to increase in tensile properties and the latter to polymer degradation. The first consideration on the use of irradiation should be if the polymer to be modified is predominantly a crosslinking or chain scission polymer. Some common chain scission polymers are polymethylmethacrylate (PMMA), polytetrafluoroethylene (PTFE), poly(vinylidene chloride), polyisobutylene (and copolymers) and polypropylene. Attempting to modify these materials with irradiation may give a surface modification, but on a degraded material. Recent high voltage accelerators with high penetration potential have been suggested as an alternative, because relatively low total dosage is required. It has also been suggested that total devices can be modified.12 The listed materials that are claimed to have been successfully modified include in particular the scission polymer PMMA. These irradiation techniques, particularly when grafting is performed in situ, require near oxygen free environments and thus create some problems for ease of manufacture. Nevertheless, irradiation is a very effective method for covalently grafting molecules to biomaterial surfaces. Plasma Plasma polymerization has been touted as the most exact method to modify biomaterial surfaces; this is in a large part due to the control over the plasma gases, and the thickness and uniformity of the modified surface. Some drawbacks to plasma are: 1. Requirement for reactor specific design based on geometry and materials of construction; 2. In almost all cases the process will be a batch operation; and 3. Contamination of the reactor and thus the requirement for cleaning cycles and for single process dedication of reactors. Several investigators have proposed simple plasma treatment to make materials more hydrophilic and thus more blood compatible.13 Plasma treatments may show initial lowering of contact angle, and will follow with an inversion within hours back to a hydrophobic surface. The latter is particularly true of elastomers. Plasma polymerized surfaces can be directly deposited on a surface, or plasma can be used to activate a surface for subsequent grafting of molecules. The latter is difficult, as the surface generally requires a short plasma discharge time, or pulsed plasma discharge. It has been hypothesized that free radicals are short lived on the surface of plasma treated materials, due largely to the rapid termination by vicinal radicals.14 A final observation on plasma-modified surfaces is that the reactivity in terms of coupling further biomolecules has been less than expected. Evaluation of coupling to plasma functionalized surfaces based on the results of XPS data suggests that a significant portion of the functional groups do not react. One possible explanation is that organic chemistry, and particularly biomolecules, react in three dimensions, and the plasma surface is highly ordered and rather two dimensional in its ability to enter into reactions.15 If one sees new biophilic surfaces as containing high

Biophilic Polymers: What’s on the Horizon?

density biomolecules such as heparin, proteins and/or growth factors, they will have to be coupled postplasma treatment and the ability to control the quantity and density of coupling may prove to be a challenge. Corona Corona treatment has been used extensively in industry for improvement of adhesion. It is also the subject of several papers on surface activation to effect grafting to surfaces.16-18 Several affinity schemes can be used to attach biomolecules to the grafted surfaces. Corona treatment is a rough process compared to plasma discharge, and takes place in an air atmosphere. This results in a highly oxidized surface, to include sufficient free radicals to effect grafting of vinyl monomers. It can be used to graft to most polymers, but does not work for most fluoropolymers. Corona and plasma can be used to make surfaces more wettable for simple adsorptive coating approaches. There are several suggested universal adsorptive coating schemes that often require a pretreatment such as corona to improve the uniformity of coating as well as the adhesive strength of the coating to the base material. CVD and PVD Chemical vapor deposition (CVD) and physical vapor deposition (PVD) have long been used in industry as barrier coatings. In the European food industry, PVD has been used to apply a layer of SiO to packaging material to prevent moisture transmission through wrapping. CVD has been used in electronics to provide moisture barriers and insulative coatings. It was hoped that these methods could be used as a base for engineering biologically active surfaces. The ability to deposit a metallic surface such as gold would give a base for self assembled surfaces.19 It has also been suggested that a ceramic based PVD surface could be activated to produce functionality (carboxyl groups) to which biomolecules could be coupled.20 Simple Coatings We define simple coatings as those that are designed to be applied by simple dipping, spraying, or by pumping solutions through devices. There is a plethora of coatings of this nature offered for short term blood compatibility. All of these coatings have performance that is material dependent; this problem is often overcome by a pretreatment such as corona. Special attention should be paid to devices made from materials that undergo bending and flexing, as these coatings may have a tendency to crack. Dislodgement of the coating may cause a problem, depending on the application. Most manufacturing engineers of biomedical devices will require a battery of tests to assure integrity of the coating, especially if the device is to be a permanent implant such as a vascular graft. While the list of short term coatings is long, the list for permanent implants is very limited. Entanglement In the late 70s and early 80s the wide acceptance of Biomer as a blood compatible material led to several appli-

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cations where the device was simply coated with the polyurethane polymer. This was best effected by finding a solvent that could attack the surface of the material to be coated and result in a mixing of polymer chains that would give a mechanical bond upon solvent removal. The same method has been used to imbrue drugs into polymers using solvents that could swell the polymer, taking in the drug and leaving it upon solvent evaporation. The latter method has been used for introducing antimicrobial agents to polymers. The same principle applies for SMAs or blooming agents. Addition of amphipathic molecules to polymers can result in the more hydrophilic portion being expressed on the polymer surface, anchored by the more hydrophobic end entangled in the polymer surface. In the case of polyurethanes, most contain processing waxes that migrate to the surface and prevent the polymer from being tacky during extrusion. One such additive, ethylene bis-stearamide wax, has been reported to be in part responsible for improved blood compatibility. Studies in our laboratory of such additives reveal that performance in blood is optimal immediately after thermal processing, but changes with time. XPS shows considerable increase over time in amide functionality, which may be less effective than the long fatty chain in improving blood compatibility. It appears that perhaps entanglement or SMAs may be difficult to control at the surface, and thus present some challenges to be solved. Texturing Professor Andreas von Recum has been one of the leading researchers in the field of tissue response to textured surfaces. He has demonstrated that textured implants with internodal distances of approximately 1-3 microns appear to have much thinner capsules, indicating less fibrotic response.38,39 One of the hypotheses offered in explanation is that the textured surface allows for more stable anchorage of fibroblasts, and less destruction of these cells due to microvibration that can lead to further inflammatory response. A possible further explanation is that the microtextrue allows for an improved production and attachment of extracellular matrix on the material surface. Research on immunoisolation membranes by Becton Dickinson presented at an ACS workshop indicated that optimal surface texture could result in minimal capsule formation and vascular structures present very close to the material surface. The latter was independent of the material of construction for the membrane. Applying microtextrue to devices is a developing technology, and one promising approach is the use of photopolymerization. This technique holds out the promise of introducing texture and chemistry to surfaces, plus the ability to use photo resists or screening techniques to pattern the sights for cellular attachment to surfaces. Application of this technology to vascular grafts will be more complicated than simply applying this technique to the inside of a polymeric tube that has laser drilled holes for porosity if one hopes to maintain other important features for grafts such as porosity for ingrowth, kink resistance, compliance, and suture characteristics.

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Ion Beam Implantation Ion beam has been used to impart surface texture in addition to surface chemistry. In the orthopedic area it is useful in improving fixation of prostheses to tissue. The Spire Corporation offers ion beam treated surfaces for improved slip properties and antimicrobial surfaces, and has made claims to improved blood compatibility.24 As with PVD, a metallic surface could provide advantages for further surface modification using self assembling techniques. Ion beam implantation can result in crosslinking the surface, and impart mechanical properties to the material. The tendency for cracking of the surface on elastomers and materials that will see chronic flexing should be considered. For elastomeric materials, and in particular silicone, the resultant surface may still present the elastomer chemistry primarily.

Chemical Methods Oxidation, Reduction, Hydrolysis Most synthetic polymers do not possess functionality to perform coupling reactions. Several of the physical methods mentioned can oxidize the surface and produce radicals, hydroxyl, carboxyl, and carbonyl groups. This oxidation can also be induced by chemical means such as acid etching or treatment with peroxides.25 For fluoropolymers that are oxidation resistant, reduction by exposure to sodium metal in napthalene can produce unsaturation in the polymer, and subsequent direct grafting of hydrogels can be effected.26 Hydrolysis, particularly of esters, can introduce functional groups such as carboxyls to surfaces. All these methods involve degradation of the polymer, and it is necessary to control or limit the degradation to have stable surfaces, as well as to prevent degradation of the mechanical properties of the base material. Addition/Substitution Some polymers such as polyurethane have sites that can be attacked by proton extracting bases such as sodium hydride. The reduced urethane group can then undergo alkyl substitution using alkyl halides such as octadecly iodide.27 An impressive body of work has been documented by the group of Cooper, showing improvement in blood compatibility and the ability to control the loading of alkyl groups on the surface. This method of introducing alkyl groups to the surface may provide better and more stable performance than previously mentioned methods using SMAs. Ozonization Treatment of materials by exposure to ozone has proven effective in activating surfaces for subsequent grafting reactions.28 Ozone can be applied as a gas or dissolved in water, which gives very attractive manufacturing features for devices with complex geometry and multiple materials, as ozone is effective on almost all polymers save fluoropolymers. Ozone is primarily an oxidative approach, and can produce peroxy radicals, but also can introduce functionality from oxidative degradation. Because it can result in degradation, care needs to be taken to control the exposure time, concentration of ozone, and temperature. These

Tissue Engineering of Prosthetic Vascular Grafts

can be done fairly easy in aqueous solutions. The ozone penetrates materials at different rates and this can be measured using colorimetric methods.29 Materials modified by ozone may need final reduction to assure that radicals are not present, and should undergo rigorous mechanical testing to assure no loss in durability of the native polymer substrate. Silanization Silanization is most frequently used for coupling to metal surfaces. The use of organofunctional silanes allows the addition of reactive organofunctional groups to the metal oxide surface. These groups can the be used to directly couple biomolecules. The most common mistake made with silanization treatments is improper application and cure methods. Excessive coating with silanes results in weak surfaces that will delaminate, and failure to properly hydrolyze the silane before drying will fail to create enough reactive silanol groups to couple during the dehydration and final formation of the polysiloxane network. Covalently attached heparin via silane activation has been released for commercial application to intravascular stents in Europe.30 Photocoupling In 1969 Knowles demonstrated the effectiveness of 2-nitro-4-azidobenzene in forming a nitrene radical upon exposure to UV that could extract a proton directly from a polymeric backbone and substitute the organo group to which it was attached directly to the polymer.31 Application of this method to the biological field was pioneered by Patrick Guire,32 leading eventually to commercialization by the BSI (Biometric Systems Inc.) Corporation.33 Numerous applications are envisioned by BSI to include enhancement of endothelial cell attachment, blood compatible surfaces, and infection resistant surfaces. While photocoupling appears to be a facile process of direct coupling to polymers, it is also material dependent, and often requires a pretreatment of the surface to remove contaminants or promote better wetting of the coupling reagents. It also often requires repeated coatings and light exposures to effect a uniform and dense coupling. This latter process can result in crosslinking of photobiomolecules within a matrix and somewhat diminish their activity. Careful cleaning of these surfaces is required to remove any unreacted or leachable species, as is the case with almost all surface modification approaches. Nevertheless, this method is receiving very favorable attention as a commercially feasible method to modify biomaterials, especially for short term use. Grafting Direct grafting of hydrogels to surfaces is only possible on a limited number of materials. The best know material is polyurethane, where the carbamate group can be activated using ceric ion initiation.34 Another material also shown many years ago to undergo direct grafting is dialysis membrane.35 Early studies on hydrogel grafted surfaces showed promises of being cell friendly.36 Ratner and Hoffman, using a baboon shunt model, later showed that hydrogel surfaces may produce emboli and, while not being thromboadherent, were in fact capable of being thrombo-

Biophilic Polymers: What’s on the Horizon?

genic. Attempts to utilize the positive aspects of grafted hydrogels by coupling bioactive agents is now commonplace. A possible advantage of this approach is lower protein adsorption and nonspecific cell adhesion. Grafted hydrogels do not require crosslinking and when coupled with biomolecules can express more bioactivity of the attached molecule. For most grafted hydrogels some sort of activation of the base polymer is required, such as plasma, corona, irradiation, or ozonization. Depending on the grafted species, additional functionalization may be required to provide optimal control of coupling with respect to loading and spacing chemistry. A good stable grafted hydrogel allows for numerous coupling strategies available from affinity chromatography techniques. Some affinity schemes have unstable bonds that are allowable for chromatographic application, but not desirable for biological application. The coupling scheme should be tested rigorously to assure that there is no leaching. In our laboratory we have found that grafting followed by intermediate functionalization allows for easier control and clean up of the grafted surface, as well as optimal loading of biomolecules, even to the point of presenting the biomolecule in high purity at the uppermost layers of the hydrogel.37

Summary The horizon for biophilic polymers was discussed in this chapter with an obvious bias toward surface modification of polymers to make them biointeractive. Our efforts have led to covalent grafted surfaces on almost all common biomaterials including metals, and stable covalent biomolecules coupled to these surfaces. In our opinion, the closest thing to a biophilic commercially available polymer in the near future would be one that maintains high mechanical performance characteristics of existing polymers and contains pendant reactive groups that can have biomolecules coupled to these functionalities in an economic process. It is doubtful that a biomolecule could be reacted into a material and still maintain its activity through processing. It would be also difficult to expect optimal presentation of the biomolecule to the blood or tissue. It may be possible for certain molecules or drugs that are targeted for release, but probably not for scaffolding applications. It is hoped that this chapter gives some useful information regarding several of the methods employed to make surfaces more biocompatible, and some appreciation for problems and opportunities that are attendant upon these methods. Knowledge of several techniques can allow for a number of combined processes that can bring optimal performance as well as economic reality to creating biocompatible surfaces. References 1. Sawyer PN. Surface charge and thrombosis. Ann NY Acad Sci 1984; 416: 561-584. 2. Chin TH, Nyilas E, Turcotte LR. Microcalorimetric and electrophoretic studies of protein sorption. Trans Am SocArtif Intern Organs 1978; 24:389-402. 3. Andrade JD. Interfacial phenomena and biomaterials. Med Instrum 1976; 7:110-120.

487 4. Ratner BD, Weathersby P, Hoffman AS, Kelly MA, Scharpen LH. Radiation-grafted hydrogels for biomaterial applications as studied by the ESCA technique. J Appl Polym Sci 1978; 22:643-664. 5. Merrill EW, Salzman EW. Polyethylene oxide as a biomaterial. ASAIO Journal 1983; 6:60-64. 6. Munro MS, Eberhart RC, Make NJ, Brink BE, Fry WJ. Thromboresistant alkyl-derivatized polyurethanes. J Am Soc Artif Int Organs 1982; 6:65-75. 7. Leininger RI. Polymers as surgical implants. CRC Critical Reviews in Bioengineering 1972; 1:333-381. 8. Gott VL, Koepke DE, Daggett RL, Zarnstorff W, Young WP. The coating of intravascular plastic prostheses with colloidal graphite. Surgery 1961; 50:382. 9. Larm O, Larsson R, Olsonn P. A new nonthrombogenic surface prepared by selective covalent binding of heparin via a modified reducing terminal residue. Biomat Med Dev Art Org 1983; 11:161-173. 10. Voorhees AB, Jaretski A, Blakemore AH. The use of tubes constructed from Vinyon “N” cloth in bridging arterial defects. Ann Surg 1952; 135:332. 11. Cahalan PT. Brite EuRam Final Technical Report, Contract Number: BREu-336, Project Number BE-597292.Title: Surface modification of Biomaterials for Biomedical Devices. European research subsidy granted in Nove. 1992 running until Nov. 1995. Report available through European Commission, Director General XII, Science, Research, and Development, Rue de la Loi 200, B-1049 Brussels, Wetstraat 200, Brussels, Belgium Office: M075 1/5. 12. Goldberg et al. Surface modified surgical instruments, devices, implants, contact lenses and the like. U.S. Patent 5,100,689, Assignee: University of Florida, Gainesville March 31, 1992. 13. Andrade JD, Triolo PM. Surface modification and evaluation of some commonly used catheter materials. I. J Biomed Mater Res 1983; 17:129-247. 14. Suzuke M, Kishida A, Iwata H, Ikada Y. Graft copolymerization of acrylamide onto a polyethylene surface pretreated with a glow discharge. Macromolecules 1986; 19:1804-1808. 15. Dias AJ, McCarthy TJ. Synthesis of a two-dimensional array of organic functional groups: Surface-selective modifications of poly(vinylidene). Macromolecules 1984; 17:2529-2531. 16. Okada T, Ikada Y. Tissue reactions to subcutaneously implanted, surface modified silicones. J Biomed Mater Res 1993; 27:1509-1518. 17. Okada T, Ikada Y. In vitro and in vivo digestion of collagen covalently immobilized onto the silicone surface. J Biomed Mater Res 1992; 26:1569-1581. 18. Okada T, Tamada Y, Ikada Y. Surface modification of silicone for tissue adhesion. Biomaterials and Clinical Applications, 1987. 19. Whiteside et al. Langmuir, 1988; 4:365. 20. Gauckler L. Personal communication at The Monte Verita Conference on Biocompatible Materials Systems. Ascona, Switzerland, 1993: October 11. 21. Csomor K, Karpati E, Nagy M, Gyorgyi-Edeleny J, Machovich R. Blood coagulation is inhibited by sulphated copolymers of vinyl alcohol and acrylic acid under in vitro as well as in vivo conditions. Thrombosis Research, 1994; 74(4):389-398.

488 22. van Boeckel CAA. From heparin to a synthetic drug: A multi-disciplinary approach. Trends in Receptor Research. Elsevier Publishers B.V. 1993. 23. Ishihara K, Nakabayashi N, Nishida K, Sakakida M, Shchiri M. Designing biocompatible materials. Chemtech, 1993. 24. Sioshansi P. Ion beam modification of materials for biomedical application. Seminar: Biomaterials: Medical and Pharmaceutical Applications, sponsored by Technomic Publishing Company, Inc, 851 New Holland Ave., Lancaster, PA: 1990. 25. Larsson N, Senius P, Eriksson JC, Maripuu R, Lindberg B. J Colloid Interface Sci, 1982; 90:127-136. 26. Yun JK, DeFife K, Colton E, Stack S, Azeez A, Cahalan L, Verhoeven M, Cahalan PL, Anderson JM. Human monocyte/macrophage adhesion and cytokine production on surface-modified poly(tetrafluoroethylene/ hexafluoropropylene) polymers with and without protein preadsorption. J Biomed Mater Res, 1995; 29:257-268. 27. Pitt WG, Cooper SL. Albumin adsorption on alkyl chain derivatized polyurethanes: I. The effect of C-18 alkylation. J Biomed Mater Res 1988; 22:359-382. 28. Yamauchi J, Yamaoka A, Ikemoto K, Matsui T. J Appl Polym Sci 1991; 43:1197-1203. 29. Kulik E et al. Poly(ethylene glycol) enriched surface: Reduced vs delayed protein adsorption and activation. Abstract from Firth World Biomaterials Congress. Toronto, Canada: 1996:May 29-June 2.

Tissue Engineering of Prosthetic Vascular Grafts 30. Cahalan L et al. Biocompatible medical article and method. U.S. Patent 5,607,475. Assignee: Medtronic Inc. 1997:March 4. 31. Knowles JR. Accounts of Chemical Research 1972; 5:90-119. 32. Guire P, Fliger D, Hodgson J. Photochemical coupling of enzymes to mammalian cells. Pharmacological Research Communications, 1977; 9(2). 33. Guire PE. Binding reagents and methods. U.S. Patent 4,722,906, Assignee: Bio-Metric Systems Inc. 1988:Feb. 2. 34. Annual Report (July 1, 1971-June 30, 1972), Medical Devices Applications Program of the National Heart & Lung Institute, Bethesda, MD. 35. Luttinger M, Cooper CW. J Biomed Mater Res 1967; 1:67. 36. Ratner BD, Horbett T, Hoffman AS. Cell adhesion to polymeric materials: Implications with respect to biocompatibility. J Biomed Mater Res, 1975; 9:407-422. 37. Hendriks M. A study on the covalent surface-immobilization of collagen. Development of biomaterials with enhanced infection resistance; a surface modification approach, Ph.D. Thesis from University of Eindhoven, 1966. 38. von Recum AF, Park JDB. Permanent percutaneous devices, CRC Crit Rev Bioeng 1981; 5:37-77. 39. von Recum AF, van Kooten TG. The influence of microtopography on cellular response and the implications for silicone implants. J Biomater Sci Polymer Edn 1995; 7:181-198.

Scaffold Engineering Material Axpects

CHAPTER 45 Bioresorbable Grafts: A Counterintuitive Approach David Fox, David A. Vorp, Howard P. Greisler

Overview

T

his chapter reviews the use of bioresorbable materials in vascular grafting. First, the theo retical basis for the use of bioresorbable materials is presented. Next, the various materials and the results of experimental work with them are discussed. Bioresorbable materials have been incorporated into vascular prostheses in a number of novel ways. Accordingly, the chapter is organized primarily by the manner in which bioresorbable materials have been incorporated. These include the bioresorbable material as a stand-alone graft, in combination with nonresorbable materials as a partially resorbable bi-component, or compound graft, and finally as a supportive scaffold over a biological graft. The chapter also discusses the use of biological grafts that have been chemically modified so as to be at least partially resorbable.

Definitions The terms biodegradable, bioresorbable and bioabsorbable have been used relatively interchangeably in the literature reporting this heterogeneous class of biomaterials. An attempt was made to reach a consensus on precise definitions at The First International Scientific Consensus Workshop on Degradable Materials held in Toronto in 1989. Biodegradation was defined as “loss of a property (of a biomaterial) caused by a biological agent.” The Workshop was unable, however, to agree on definitions of degradation, bioabsorption and bioresorption.1 Therefore, in this chapter the terms biodegradable, bioresorbable and bioabsorbable are in general applied in accordance with the original authors, without implication of specific mechanisms of degradation.

Introduction Despite the successful application of prosthetic grafts in the replacement of large arteries, the performance of prosthetics as medium and small caliber replacements has been less than satisfactory, resulting in both early and late graft failures, increasing the need for reoperation and limb loss. Factors contributing to these failures include thrombogenicity, inadequate tissue ingrowth and hyperproliferation of tissue with deposition of extracellular matrix resulting in myointimal hyperplasia.2 Biodegradable grafts have been investigated as a potential alternative to conventional biostable prosthetic grafts. The concept of a biodegradable graft emerged from the work of Arthur Voorhees’ group at Columbia University. In 1954 Voorhees reported the first truly successful clinical application of prosthetic arterial prostheses, replacing 17 abdominal aortas and 1 popliteal aneurysm with grafts composed of Vinyon-N cloth tubes.3 The success of Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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the Vinyon-N fabric prostheses was in large part a function of its porosity, which allowed the ingress of capillaries and fibroblasts. This granulation tissue served as the nidus for the formation of an organized inner surface of flattened cells.4 Thus, it became apparent that the graft material itself may be needed only transiently, to function as a scaffolding for regeneration of the vascular wall rather than as a permanent conduit. After Voorhees’ success, many other materials were proposed. Sigmund Wesolowski and his co-workers, working at Walter Reed, evaluated 45 prospective materials and confirmed the concept that the porosity of a graft material and not its inertness determined the success of a prosthetic vascular graft.5 Prior to this time, the prevailing concept was that the biological inertness of a prosthetic material was the critical determinant of its clinical success as a vascular graft.6 Wesolowski determined that grafts constructed of materials of relatively high porosity were resistant to late calcification and occlusion. The problem of hemorrhage during implantation, however, remained an important obstacle. To overcome this problem, the concept of constructing a composite graft emerged. The ideal composite graft would exhibit the property of low porosity at implantation, by virtue of a degradable component combined with a permanent component of high porosity. Over time, the degradable component would be depleted, leaving a high porosity graft capable of being incorporated by tissue.7 The feasibility of incorporating absorbable collagen and gelatin components into Dacron grafts had been reported by Humphries and Bascon in 1961.8,9 Weselowski took a different approach, suggesting the use of a temporary scaffold fabricated from a slowly absorbable polymeric material as a vascular graft.5 This scaffold would enable the arterial wall to reorganize by virtue of natural repair processes.7 The American physiologist Claude Guthrie has been credited with the origin of this concept. In 1919 he wrote, “To restore and maintain mechanical function an implanted segment only temporarily restores mechanical continuity and serves as a scaffolding or bridge for the laying down of an ingrowth of tissue derived from the host.”10 Since Wesolowski’s report, many additional biodegradable materials have been proposed for use as components in the fabrication of vascular prostheses. These materials have been incorporated into grafts in a number of interesting and novel ways that will be described in this chapter. The majority of investigators have used the biodegradable material as a stand-alone graft. Others have constructed compound grafts wherein the degradable material is combined with a nonabsorbable prosthetic or biological graft. The degradable materials have been applied as a lining or wrap and in some cases have been interweaved with the prosthetic. The biomaterials used in the fabrication of bioresorbable vascular grafts have a unique dual function. They serve as temporary vascular conduits while simultaneously inducing the regeneration of the arterial wall. As will be elaborated, complex interactions between the bioresorbable material and host macrophages are a critical step in the process of arterial regeneration.

Tissue Engineering of Prosthetic Vascular Grafts

Single Component Resorbable Grafts The materials most commonly used in the fabrication of resorbable grafts have been synthetic polymers based upon naturally occurring hydroxy acids such as lactic and glycolic acid.11 The rate of resorption of these copolymers is determined to a degree by the ratio of lactide to glycolide rings and decreases as the percentage of lactide component increases.12 The primary mechanism of resorption is by the hydrolytic cleavage of ester bonds. Inflammatory cells such as macrophages, lymphocytes and neutrophils perform the final degradation and resorption.11 Polyglactin 910 (Vicryl) Bowald’s group from Sweden is credited with the first report of the use of a totally biodegradable prosthesis as a stand-alone replacement for large caliber vessels.7 Knitted mesh composed of Polyglactin 910 (PG910), “Vicryl,” manufactured by Ethicon Inc., was preclotted and interposed as tube and patch grafts into the thoracic aorta of growing pigs. PG910 is a polymer composed of glycolic and lactic acids and thus is absorbed by hydrolysis within 70 days.13 The mesh had a pore size of 400 x 400 microns.14 The grafts were explanted at intervals up to 6 weeks. All grafts remained patent and no instances of graft rupture or aneurysmal dilatation occurred. At 40 days only small fragments of the PG910 remained and the grafted areas displayed some similarities to normal aortic walls. The luminal surface was covered with a monolayer of mature endothelial-like cells in a mosaic pattern. Beneath this was a layer of longitudinally directed smooth muscle-like cells. No internal elastic lamina was seen, although scattered fibrillar elastic material was found interspersed amongst the smooth muscle-like cells. An outer “adventitial” layer was reconstituted as well.13 With respect to the source of the repopulating cells, Bowald observed that the endothelium and smooth muscle-like cells seemed to be derived from the cut edges of the native vessel. Occasional inflammatory and giant cells were noted around residual PG910 filaments in the 20 day explants.14 Late results after 2 years were reported for one pig that received a 10 mm by 4 cm tube graft in the thoracic aorta. The graft was patent angiographically. On gross examination it was free of thrombus, calcification or fatty infiltration. Of note, the diameter of the grafted area had increased proportionally in size with the nongrafted aorta and was not aneurysmal. They also reported reconstitution of a coarse internal elastic lamina.15 Based on these observations, Bowald sought to determine the maximal regenerative capacity of arterial tissue. To this end, the descending thoracic aorta of 20 pigs was bypassed with 8-10 mm diameter double layer Vicryl segments ranging from 7-20 cm in length. All animals receiving a graft greater than 15 cm in length died from graft rupture. There were no ruptures in the animals receiving grafts less than 15 cm and the histologic appearance of these grafts was equivalent to those in the previous reports.16 Extensive implant and in vitro work with lactic and glycolide copolymers have also been done by our group in the research laboratories of Loyola University. We have in-

Bioresorbable Grafts: A Counterintuitive Approach

vestigated PG910 and other copolymers including polyglycolic acid (PGA) and polydioxonone (PDS), as well as combinations of these polymers. PG910 grafts were implanted into the abdominal aorta of rabbits. Likewise, the transinterstitial migration and proliferation of ultrastruc-

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turally primitive mesenchymal cells that differentiated into smooth muscle-like myofibroblasts and the establishment of a confluent endothelial-like luminal lining cells was demonstrated, thus confirming Bowald’s earlier observations (Figs. 45.1A-C).14 This process was found to parallel the time

Fig. 45.1. Hematoxylin-eosin histologic sections (x25) from 1 month post implantation. (A, top) Dacron prosthesis demonstrating a confluent endothelial surface with a small subendothelial layer. (B, middle) Polyglactin 910 prosthesis demonstrating a thick subendothelial cellular inner capsule with an endothelial luminal surface. (C, bottom) Polydioxanone prosthesis demonstrating an endothelial lined luminal surface with an intermediate subendothelial layer.

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course of macrophage-mediated prosthetic dissolution. Peak inner capsule (IC) thickness was achieved at 2 months. Again, there was no graft related morbidity or mortality. Twenty percent of the PG910 grafts become aneurysmal.17 In order to further assess the potential contribution of transanastomotic pannus ingrowth to the arterial regeneration seen, prostheses were constructed of three 10 mm segments with a PG910 segment interposed between two Dacron segments. Grafts were again implanted in rabbit aortas and harvested at intervals up to 4 months. All grafts remained patent with no aneurysms or stenoses. Inner capsule thickness in the PG910 segments increased only during the interval from 2 weeks to 2 months and was statistically greater than either Dacron segment at 1 and 2 months. Inner capsules of PG910 segments at 1 month were composed predominantly of myofibroblasts. Pannus ingrowth was limited to the first 2 mm of the inner capsules of the Dacron segments, the remainder consisting of a fibrin coagulum. These findings essentially ruled out transanastomotic pannus ingrowth as the primary source of cells replacing absorbable vascular prostheses.18 Polyglycolic Acid (Dexon) In 1982, we reported the regeneration of major components of the rabbit aorta over a scaffold of absorbable porous polyglycolic acid (PGA). PGA has the simplest structure of the linear aliphatic polyesters. It is hydrophilic and loses its strength rapidly over 2-4 weeks.11 Woven PGA, manufactured by Davis & Geck, was fabricated into 3.5 mm diameter grafts and implanted into the abdominal aorta of rabbits and explanted over a seven and a half month period. Seventy-six percent (76%) of the animals were found to have maintained a lumen of 3-4 mm in the region of implantation. Eleven percent (11%) of the animals demonstrated aneurysmal dilatation in this region; however, no animal died of graft rupture. In addition, there was no evidence of perigraft hematomas or of false aneurysm formation. There were no instances of graft thrombosis or infection. Experimental regenerated aortas were subjected to bursting strength determinations and withstood hemodynamic forces at three to five times systolic pressure. Histologic analysis revealed some marked similarities to native aorta. The regenerated aortas were composed of a luminal surface containing factor VIII positive endothelial cells, a subendothelial zone containing smooth muscle-like myofibroblasts amidst dense fibrous tissue and an outer layer of vascularized connective tissue. Elastin regeneration, however, was minimal. Some lipid-laden macrophages and histiocytes were observed in the six month explants, suggesting possible early atherogenesis. Intimal hyperplasia resulting in severe graft stenosis was observed in 13% of animals. The PGA graft material underwent progressive dissolution. At two weeks the PGA was densely invaded by mesenchymal cells, histiocytes and giant cells. Complete resorption occurred between three weeks and three months. Important questions emerged from this study. Principally, it remained to be defined what were the factors responsible for the initiation and modulation of the repair process. Additionally, the origin of the cells in the regener-

Tissue Engineering of Prosthetic Vascular Grafts

ated aorta was unknown, as was the atherogenicity of the vessels.19 Stimulated by these early results, our group compared Dacron aortic prostheses to PGA. The Dacron grafts were fabricated to specifications that closely matched the physical characteristics of woven PGA, most notably in terms of pore diameter, pore area and wall thickness. Over the 12 month observation period both the PGA and Dacron grafts remained free of thrombosis, infection or hemorrhagic complications. Again, 10% of PGA grafts showed moderate aneurysmal dilatation and 15% demonstrated significant intimal hyperplasia.20 After three months the PGA, as before, was found to have been resorbed and in its place a thick, dense neointima lined with a confluent layer of endothelial cells was seen. Beneath this layer were numerous smooth muscle like myofibroblasts in a circumferential and longitudinal orientation. The inner capsules of the Dacron grafts, however, exhibited only matrix and small numbers of myofibroblasts in a radial orientation. The flow surface was thin, fibrinous and acellular. The Dacron grafts explanted at 1 year contained extensive calcium deposits and some advanced atherosclerotic plaques. The PGA explants at 1 year remained remarkably stable with the exception of an occasional lipid-laden histiocyte, as had been previously observed. Despite their similarities in mechanical porosity, these two graft materials achieved remarkably different histologic outcomes, suggesting that porosity was not the only factor active in graft healing. Insight into the pathophysiology of these differences was gained by studying the grafts explanted after two weeks. As in the previous experiment, an inflammatory reaction containing macrophages, polymorphonuclear leukocytes and giant cells was observed in the tissues surrounding the PGA grafts. Additionally, numerous macrophages were found within the mesh of the PGA grafts. The Dacron grafts as well were surrounded by an inflammatory reaction; however, cellular infiltration was limited to the outer capsule with little or no permeation of the woven Dacron. It was hypothesized that phagocytosis of PGA by macrophages inducing their phenotypic alteration and synthesis of growth factors chemotactic and mitogenic for endothelial cells and myofibroblasts, as well as subtle differences in the surface electronegativity of the materials, could have accounted for differences in transinterstitial migration of inflammatory cells.20 Additional speculations were made regarding the role of growth factors in inducing myofibroblast and endothelial cell ingrowth and deposition. Macrophages were known to stimulate the proliferation of endothelial cells, smooth muscle and fibroblasts through the elucidation of “macrophage-derived growth factors (MDGFs).”21-23 The degree of macrophage infiltration corresponded temporally with myofibroblast proliferation as quantitated by autoradiography with tritiated thymidine and suggested the possibility that a macrophage derived factor played an important role in the modulation of myofibroblast response. The possibility that the degradable materials had directly resulted in

Bioresorbable Grafts: A Counterintuitive Approach

Fig. 45.2. Top: Midportion of 1 month Dacron specimen showing thin inner capsule (IC) composed of fibrin coagulum and minimal cellular repopulation. (Hematoxylin-eosin; x50). Bottom: Midportion of 1 month polyglycolic acid specimen showing greatly thickened IC of circumferentially and longitudinally oriented smooth muscle-like myofibroblasts beneath endothelial-like flow surface. More peripheral are prosthesis and well vascularized outer capsule. (Hematoxylin-eosin; x50).

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Fig. 45.3. Midportion of PDS specimen at 2 months showing capillary invading inner capsule and communicating with luminal surface. Endothelial-like cellular luminal surface of specimen appears to be continuous with capillary wall.

Fig. 45.4. Midportion of PDS specimen at 12 months showing maintenance of smooth confluent layer of endothelial-like cells on luminal surface. (Hematoxylineosin; x390).

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macrophage stimulation was suggested by the demonstration of PGA remnants within the cytoplasm of macrophages by TEM. Additional speculations surrounded the potential roles of PDGF, ECGF (now FGF-1), and FGF-2 and of mechanical forces on the tissue bed, which were found to increase as the PGA degrades and the integrity and mechanical strength of the graft material was lost.20 Polydioxanone (Pds) The ability of lactide/glycolide polymers to induce arterial regeneration was also demonstrated with grafts fabricated from PDS, a more slowly resorbed polymer.24 Grafts, 24 x 4 mm in size and composed of woven PDS were again interposed into the abdominal aorta of rabbits. The animals were studied at intervals up to one year. Again, no aorticrelated deaths or hemorrhages occurred. Aneurysmal dilatation developed in 10% of specimens, whereas stenosis from neointimal hyperplasia occurred in 15%. As with PGA (Fig. 45.2), a vessel wall composed of myofibroblasts and matrix covered by a confluent endothelial-like flow surface resulted. The PDS implants (Figs. 45.3 and 45.4) yielded an inner capsule (IC) thickness as well as compliance and strength characteristics that were comparable to the PGA grafts. Not unexpectedly, tissue formation with the PDS grafts occurred more gradually. These grafts reached maximal IC thickness between 3 and 6 months as compared to 2 months for PGA. The difference in the rates paralleled the differing dissolution rates of the two materials.25 These observations reinforced the concept that the process of arterial regeneration was regulated by some factor or factors related to the absorption of the material by macrophages. Mechanical Stimuli of Arterial Regeneration A study by Vorp et al to evaluate the stress history within the inner capsule demonstrated that, early in the resorption phase, the neointima is under a state of circumferential compression.26 Later in the resorption phase, as the graft material weakens upon dissolution and as the neointimal tissue strengthens upon reorganization, the inner capsule stress becomes tensile. The occurrence of compressive stresses parallels the time point and location of maximal mitotic index for the neointimal tissue. We showed using the well established rabbit model that the mitotic index of inner capsule myofibroblasts in PG910 grafts was significantly higher during the early period of resorption than in the later period.27 Using a pulse duplicator apparatus developed at the University of Pittsburgh,28 changes in the dynamic compliance of PG910 grafts occurring over 0 to 36 weeks were assessed.29 Using the rabbit model, explanted PG910 grafts were exposed to realistic pulsatile hemodynamics in vitro. A rapid increase in compliance was noted after 3 weeks, leveling off at 13 weeks. In a companion study, compliance increased over time in the proximal and distal graft segments compared to baseline. Loss of compliance has been cited as an important mechanism in the failure of small and medium sized Dacron and ePTFE implants.30 Thus, by achieving greater compliance over time, resorbable grafts appear

Tissue Engineering of Prosthetic Vascular Grafts

to have real potential for improved efficacy over prosthetics in current use. Characteristics of the Healing Response In addition to rate of resorption and kinetics of IC formation, additional studies were performed in order to further characterize differences between the healing responses induced by the various materials. The prostaglandin metabolite contents of inner capsules formed in the presence of PDS, PG910 and Dacron containing grafts were compared. PDS inner capsules were found to have greater ratios of 6-keto-PGF1α:TxB2 as compared to PG910 and Dacron implants. Notably, the ratio of 6-keto-PGF1α:TxB2 for PDS was nearly identical to its normal aortic control. This favorable ratio suggested that grafts containing PDS may be less thrombogenic and again suggests the potential for clinical efficacy of resorbable grafts.31 The kinetics of collagen deposition induced by PDS, PG910 and Dacron containing grafts was determined by examining the inner capsules of grafts explanted at 1, 3 and 12 months. In the Dacron grafts, collagen formation peaked at 1 month at approximately 12 µg/mg. In contrast, the collagen content of the inner capsules of the PDS and PG910 grafts at 1 month was around 27 µg/mg. This amount of collagen is essentially equivalent to that found in normal infrarenal rabbit aorta. The level of collagen in the Dacron grafts at one year remained the same as at 1 month. The level of collagen in the PDS and PG910 grafts increased to approximately 38 µg/mg at 3 months, after which it stabilized.32 The kinetics of myofibroblast proliferation in the inner capsules of PDS, PG910 and Dacron implants was also analyzed, again in the rabbit infrarenal aorta model. Grafts were explanted at intervals up to 52 weeks. Mitotic indices were determined by autoradiography using tritiated thymidine. Of interest was the finding that the mitotic index associated with each graft material was parallel to that material’s rate of resorption. The peak mitotic index for PG910 was 28.34 in a 24 h period and occurred at 3 weeks postimplant, whereas the peak for PDS was 7.50; this occurred at 4 weeks and was prolonged. There was little evidence of inner capsule formation in the Dacron grafts until 12 weeks postimplant and the mitotic indices for Dacron grafts never exceeded 1.22.27 Macrophage/Biomaterial Interactions In order to further characterize the contribution of the macrophage/biomaterial interaction in the elaboration of factors regulating the process of arterial regeneration, a series of in vitro experiments were performed.33-36 In each series of experiments, rabbit peritoneal macrophages were harvested and cultured in Minimum Essential Medium (MEM) with platelet-poor serum and containing particles of Dacron, polyglactin 910 or no biomaterial (Fig. 45.5A,B). MEM conditioned in this way was collected and mitogenicity assays were performed with quiescent fibroblasts (BALB/c3T3), rabbit aortic smooth muscle cells, and murine capillary lung (LE-II) endothelial cells. Mitogenic ac-

Bioresorbable Grafts: A Counterintuitive Approach

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Fig. 45.5. (A, top) Phase contrast photomicrograph of peritoneal macrophages following six weeks in culture in the presence of the bioresorbable polyglactin 910. Intracytoplasmic polyglactin 910 inclusions can be seen (original magnification x400). (B, bottom) Phase contrast photomicrograph of peritoneal macrophages following five weeks in culture in the presence of Dacron. Macrophages can be seen adherent to Dacron particles but no intracytoplasmic inclusions are observed (original magnification x100).

tivity was assayed by scintillation counting of tritiated thymidine incorporation into deoxyribonucleic acid (DNA). In order to examine these processes in a model more accurately reflecting the clinical setting, macrophages were also harvested from rabbits that had been rendered hyperlipidemic through an atherosclerotic diet.33-35 All cell types grown in conditioned media from macrophages exposed to PG910 demonstrated significantly greater increases in tritiated thymidine incorporation as compared to those grown in media conditioned from mac-

rophages exposed either to Dacron or to no biomaterial. In fact, it could be demonstrated that macrophages harvested from rabbits fed a cholesterol rich atherogenic diet exposed in culture to Dacron produced conditioned media that exerted an inhibitory effect on endothelial cell thymidine incorporation. In order to isolate the factor or factors responsible for these alterations in mitogenic activity, Western blot analysis was performed using antibodies raised against PDGF-B chain, acidic fibroblast growth factor (FGF-1) and basic fibroblast growth factor (FGF-2). Only anti-bFGF

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Tissue Engineering of Prosthetic Vascular Grafts

Fig. 45.6. (A) Rabbit aortic smooth muscle cell bioassay results. Conditioned media from macrophages harvested from New Zealand White rabbits fed a normal diet and cultured from 1 through 7 weeks in the presence of either no biomaterial (first bar), polyglactin 910 (second bar), or Dacron (third bar). Results are expressed as counts per minute of the sample divided by counts per minute of the fetal bovine serum positive control x 100 (mean ± SD). (B) Rabbit aortic smooth muscle cell bioassay results. Conditioned media from macrophages harvested from New Zealand White rabbits fed a 2% cholesterol/6% peanut oil diet and cultured from 1 through 7 weeks in the presence of either no biomaterial (first bar), polyglactin 910 (second bar), or Dacron (third bar). Results are expressed as counts per minute of the sample divided by counts per minute of the fetal bovine serum positive control x 100 (mean ± SD).

antibody immunoreacted with samples of the conditioned media.33 In order to further define the contribution of bFGF to the mitogenic response, conditioned media was preincubated with neutralizing anti-basic-FGF antibody. In rabbits fed normal diets, Dacron induced mitogen release was completely inhibited by anti-basic-FGF antibody. PG910 mitogen release was diminished, but to a significantly less degree (36%). Thus it appeared that much of the mitogenic effect on smooth muscle cell was mediated via basic-FGF.35 In another series of experiments, the effect of conditioned media on rabbit aortic smooth muscle cells was assessed. Smooth muscle cells grown in conditioned media from macrophages exposed to PG910 again resulted in significantly higher tritiated thymidine incorporation than.

smooth muscle cells grown in conditioned media from macrophages exposed to Dacron (Fig. 45.6A). When macrophages were obtained from atherosclerotic rabbits a generalized decrease in the incorporation of tritiated thymidine by smooth muscle cells occurred in both the Dacron and PG910 groups. The macrophages from atherosclerotic rabbits cultured with Dacron, however, showed a particular increase in the mitogenic factors most active on the smooth muscle cells (Fig. 45.6B). To further characterize this effect, Western blot and immunoprecipitation studies were performed, and as before, the presence of FGF-2 was demonstrated, whereas PDGF and FGF-1 were not. TGF-β inhibitory activity was demonstrated to be increased for the atherosclerotic macrophage containing media. These findings support the hypothesis that interactions between Dacron,

Bioresorbable Grafts: A Counterintuitive Approach

macrophage and smooth muscle cells play a role in the pathogenesis of graft failure.34 Polyurethane/Poly-L-Lactide Copolymers (Pu-Plla) A different approach has been taken by the polymer chemists at the University of Groningen in The Netherlands. They have worked extensively with copolymers of polyurethane and poly-L-lactide (PU-PLLA) with the goal of combining the properties of the two materials to produce an antithrombogenic, flexible as well as resorbable material. PUPLLA was described by Gogolewski in 1982.37 Lommen et al interposed 1.5 x 10 mm PU-PLLA grafts into the abdominal aorta of rats.2 These grafts were of 40-50 µm porosity and were explanted at 3, 6 and 12 weeks. These grafts were compared to a control group that received PTFE grafts of the same dimensions and of 30 µm porosity. At explant all grafts were patent and there was no evidence of aneurysm formation, dilatation or other complications in either group. By 6 weeks the PTFE grafts had developed a complete “endothelial” lining that was continuous with that of the adjacent vessel. However, this lining detached easily during processing, suggesting that there was poor attachment with the prosthesis. The internodal spaces of the graft appeared to contain fibrin. Similar findings were noted in the 12 week PTFE explants. The PU-PLLA grafts demonstrated arterial regeneration. A complete “endothelial” lining was observed at three weeks. Likewise, the outer surface of the grafts was covered with vascularized connective tissue that was morphologically indistinct from native adventitia. At 6 weeks giant cell activity was observed within the graft walls, most prominently at the periphery. By 6-12 weeks a subintimal layer containing “smooth muscle cells,” collagen and elastin (demonstrated by Orcein staining) had developed. Multiple fenestrated elastic laminae were observed as well. The body of the PU-PLLA graft gradually became infiltrated with cells, and this process was associated with progressive degradation of the graft substance.2 Having demonstrated that biodegradable PU-PLLA stimulates the regeneration of an arterial wall they next sought to optimize some of the properties of the grafts responsible for this process. They postulated that the compliance of the grafts was an important property and they performed in vitro stress/strain measurements of grafts composed of several combinations of PU and PLLA and compared them to native artery. They determined that grafts composed of 5-20% PLLA possessed compliance before implantation comparable to rat aorta. 38 They subsequently investigated the effect of varying the pore size of grafts composed of 95%/5% PU/PLLA in the rat model. They determined that there was no distinct advantage of a pore size of 40 µm as compared to pores in a gradient of 10-100 µm from the inner to the outer regions of the graft in terms of arterial regeneration. The advocated the use of a pore gradient, however, as the smaller pores on the luminal surface were thought to decrease thrombogenicity.39 They compared these compliant grafts to less compliant grafts composed of biostable PU and of PU/PLLA surrounded by PTFE. Partial arterial regeneration was seen in

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all groups; however, only the PU/PLLA grafts reconstituted an elastic lamina. They concluded that both biodegradation and compliance were important stimuli of arterial regeneration.40 To characterize the functional characteristics of the regenerated arteries, endothelialization and prostacyclin synthesis were determined using a bioassay and RIA for the prostacyclin metabolite 6-oxo-PGF1α. Comparisons were made between PU/PLLA and PTFE grafts and normal endothelium. Looking at prostacyclin production per unit graft area covered with neoendothelium, there were no significant differences between the two graft types and native artery. Endothelialization and healing, however, as determined by light and electron microscopy were incomplete in the PTFE grafts, even after 12 weeks.41 Additional ultrastructural analyses were performed and, as was observed by our group, epithelioid cells and multinucleated giant cells were seen to engulf polymer particles of the disintegrating grafts.42 The long term fate of PU/PLLA grafts was assessed in a group of 8 implants. After 1 year all grafts were found to be patent. Three of the grafts were of normal shape and were visibly pulsatile. Two grafts were mildly dilated, 2 were aneurysmal and 1 had sustained a dissection. These grafts were not visibly pulsatile. Histologic analysis of the “neomedia” revealed circularly arranged smooth muscle cells (as in normal artery) in the normally shaped implants. However, in the nonpulsatile implants, the neomedia had a predominantly longitudinal arrangement. They concluded that the pattern of arrangement of smooth muscle cells in the neomedia affects the fate of the regenerated arteries.43 There was uncertainty as to the origin of the cells in the regenerated arteries. Hypothesized origins have included migration from the cut ends of the adjacent vessel and the granulation tissue surrounding the graft, as well as deposition of circulating cells onto the luminal surface.44 To gain insight into the source origin of the cells in the arteries regenerated from PU/PLLA, sequential explants at intervals of 1 h to twelve weeks were performed. They observed that the endothelium originated from adjacent intima. The smooth muscle cells seemed to originate from the media of the adjacent aorta; however, they could not exclude the other postulated mechanisms which would be required for clinical applicability, and proposed experiments with longer grafts to decrease the likelihood of trans-anastomotic cell migration.44 Other work with lactide polymers has been done by Hanson’s group at the University of Texas. They fabricated grafts with copolymers of L-lactide, DL-lactide, and ε-caprolactone in various compositions. Lactide polymers were chosen for their strength, ε-caprolactone for their flexibility. The grafts were 3 mm by 8 cm, nonwoven, nonporous and completely resorbable. The mechanical properties and degradation rates of the grafts were evaluated in vitro. Circumferential tension, compliance, and kink angle were determined. Tensile strength and modulus of elasticity were found to be in the range of normal artery; however, the grafts were significantly less compliant than artery. It was anticipated that the addition of porosity, a necessity for in vivo

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studies, would improve compliance. Further investigations were planned.45 Polyurethane Elastomers (PU) Extensive work has been done in the evaluation of small diameter vascular grafts composed of polyurethane elastomers. Polyurethanes hold promise by virtue of their superior compliance as compared to conventional prosthetics. By and large, these polyurethanes are nonresorbable (biostable).46 Gogolewski fabricated vascular grafts from a bioresorbable segmented aliphatic polyurethane (PU). Like the PU/PLLA grafts, these were degradable, microporous and compliant. Grafts of 1.5 mm inner diameter and 0.7 mm wall thickness were interposed into 13 mm defects in rat abdominal aorta and explanted at intervals between 3 and 16 weeks. The luminal surface of the graft consisted of interconnected pores of 5-10 µm. The outer surface of the graft had a porosity of 20-50 µm. The tissue response was reported as being comparable to that previously observed with polylactide and polyglycolide. Again, prosthetic debris were surrounded by giant cells and macrophages. A “glistening neointima” was observed at 3 weeks, with “no tendency” towards intimal thickening. Smooth muscle cells and elastin fibers were observed in the intima. By four months almost complete degradation of the PU had occurred. Stress/strain analysis revealed mechanical characteristics similar to native aorta.46 Encouraged by these results, 8 mm by 6 cm grafts were implanted into the infrarenal aorta of 50 young pigs. The grafts were explanted at intervals of 1 month to 1 year. Eighteen pigs died as a result of prosthetic rupture, infection or thrombosis. Ten pigs received no aspirin and a regular diet. These grafts were all thrombosed by 90 days. Twenty-five pigs received aspirin; 55% of these grafts remained patent at 90 days. Only 3 were patent at 1 year. Of 15 pigs receiving a diet high in lipid, 7 remained patent at 1 year. Three of these animals were noted to have dilation of the prosthesis similar to that observed by our group in PG910 grafts placed in the abdominal aorta of rabbits fed an atherogenic diet. 47 It was postulated that the enhanced patency of these grafts was related to an anti-thrombotic effect of the lipid-rich diet.48 Histologic analysis again revealed macrophage and giant cell infiltration of the prosthesis and phagocytosis of prosthetic debris. A neointima and neomedia containing longitudinal and transverse smooth muscle cells developed by 180 days. A fibrous neoadventitia was observed at early time points; however, at 1 year this was comprised of fatty tissue. Elastic fibers dispersed throughout the muscle layer were observed. The prostheses were completely resorbed by 1 year. Prostacycline synthesis was found to be comparable to normal aorta with no apparent differences between groups.48,49

Preventing Aneurysmal Dilatation The chief obstacle in the development of a clinically useful resorbable graft has been the potential for loss of mechanical and structural integrity in the graft prior to the regeneration of an arterial wall. Adding to this concern is the finding that when evaluated in hyperlipidemic models,

Tissue Engineering of Prosthetic Vascular Grafts

which more closely approximate the situation of clinical vascular grafting, resorbable grafts have had an increased incidence of graft dilation, aneurysm formation and rupture.47-49 Investigators have taken several approaches to counteracting this problem, including incorporating resorbables with enhanced resistance to degradation,7,50-56 increasing the rapidity of tissue regeneration with growth factors57,58 and cell seeding.59,60 Others have taken an intermediate approach and constructed compound or partially resorbable vascular prostheses combining resorbable and permanent materials. Enhancing the Resistance to Degradation of Resorbable Materials Our group evaluated woven bicomponent totally bioresorbable grafts fabricated from compound yarns containing 74% PG910 with 26% polydioxanone (PDS). The resorption kinetics of PDS are slower than PG910. Thirtyseven grafts were implanted into the infrarenal aorta of rabbits and explanted at intervals between 2 weeks to 12 months. Complete resorbtion of the PG910 occurred by 2 months and the PDS by 6 months. All grafts were patent at explant. One graft developed a 50% stenosis that appeared to be thrombotic in origin. No aneurysmal dilatation was seen in any graft. Inner capsule thickness was intermediate to that seen with PDS and PG910.51 The Groningen group hypothesized that the dilatation observed with their PU/PLLA grafts was related to technical problems with the dip-coating process of graft preparation.52 They produced a modified 2-ply graft with a more biostable inner layer. The inner layer was made highly antithrombogenic by crosslinking of a mixture of linoleic acid and an aliphatic polyetherurethane with dicumyl peroxide. The outer ply was a (95/5) mixture of polyesterurethane and poly(L-lactide). Twenty-two two-ply grafts were implanted in the abdominal aorta of rats. All grafts remained patent at least up to 1 year and did not exhibit any aneurysm formation. As in the earlier formulations, the inner layer was covered with endothelial cells and several layers of smooth muscle cells.53 Galletti’s group at Brown University approached the problem of aneurysmal degeneration of Vicryl grafts by coating the yarns used in the fabrication process with slowly resorbed polyesters, thus extending their functional life. They produced tubular conduits, 1 mm by 3 mm, composed of triple ply Vicryl coated with a 1:1 composite of poly-DLlactic acid and poly-2,3-butylene malate or poly-2,3-butylene fumarate. Coated or uncoated Vicryl grafts were implanted subcutaneously in mice. Samples were retrieved at intervals between 2 and 12 weeks and the remaining mass of polymer implant was analyzed for the presence of Vicryl using gel permeation chromatography. After 8 weeks Vicryl was undetectable in the uncoated implants. Grafts coated with poly-2,3-butylene malate or poly-2,3-butylene fumarate, however, retained 48 and 88% of their Vicryl content respectively.54 Galletti’s group next constructed 8 mm by 13 cm grafts of triple ply Vicryl coated with a 1:1 composite of poly-DLlactic acid and poly-2,3-butylene fumarate. To minimize oozing through the graft interstices, they added a fourth

Bioresorbable Grafts: A Counterintuitive Approach

outer layer of a 1 cm wide ribbon of Vicryl mesh wrapped in a spiral around the graft. The grafts were inserted in the infrarenal aorta of dogs. Grafts were retrieved at intervals up to 12 weeks. They found that at 8 weeks the polymer was substantially reduced and by 12 weeks had nearly disappeared. Stress/strain analysis of the resulting inner capsules demonstrated properties similar to natural vessels.55 They next investigated prostheses composed of Vicryl coated with either poly-DL-lactic acid and poly-2,3-butylene fumarate (PDLA-PBF) or poly-DL-lactic acid and poly2,3-butylene succinate (PDLA-PBS). Grafts (8-9 mm in internal diameter, 8-10 cm long) were implanted in the infrarenal aortic position of dogs and explanted at intervals up to 24 weeks. All animals receiving coated prostheses survived, whereas one animal receiving an uncoated Vicryl prosthesis died from graft rupture. Grafts were patent in 14 of 18 animals at the time of retrieval. Histologic analysis revealed no cellular invasion of the Vicryl mesh in the PDLA-PBF coated grafts. The PDLA-PBS coated grafts, however, demonstrated cellular elements, and fusion of inner and outer tissue layers could be demonstrated. Isolated patches of endothelium-like cells were noted in the mid-region of the PDLA-PBS coated grafts; otherwise, composition of the inner and outer capsules did not differ significantly between groups.56 Increasing the Rapidity of Tissue Regeneration with Growth Factors Our group sought to minimize the likelihood of graft rupture by accelerating the regeneration of the arterial wall through the local application of growth factors. We documented our ability to attach FGF-1, a potent endothelial cell mitogen, and FGF-1-heparin complexes to PDS prosthesis.58 Dacron or PDS grafts were then interposed into the infrarenal aorta of rabbits and explanted at intervals up to 30 days. We confirmed the method of attachment of FGF-1 to biomaterial surfaces and documented the retention of FGF-1 to the graft while in circulation. At 1 month all grafts were patent and without stenoses or aneurysm formation. No significant differences were demonstrated between FGF-1 and control groups in terms of cellularity of luminal surfaces, inner capsule thickness or capillarization of the inner capsule.61 After explantation, residual [125I]FGF-1 was eluted from the prostheses. Intact FGF-1 was identified by SDS gel electrophoresis. Similarly prepared prostheses were explanted after 7 days and their FGF-1 eluted off for in vitro documentation of activity. Eluted FGF-1 was found to have retained its mitogenic properties, causing a 1000-1200% increase in [3H]thymidine incorporation into newly synthesized DNA in test murine LE-II cells.61 We hypothesize that the lack of difference in healing seen between control and growth factor treated grafts relates to the inactivity of bound FGF-1. It is also possible that differences may have occurred earlier than the 30 day explant time used in this study. We have subsequently altered our approach and are applying growth factors to grafts in polymerized suspensions of fibrin glue.62 We have not yet applied this technique to resorbable grafts.

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Cell Seeding The Groningen group sought to accelerate the regeneration of the neomedia through the application of smooth muscle cell seeding.59,60 Cultured smooth muscle was seeded into PU/PLLA biodegradable vascular grafts and implanted into the abdominal aorta of rats. Arterial wall regeneration was evaluated at intervals between 2 h and 1 week. The seeded grafts showed a more rapid and uniform development of the neomedia (clearly discernible at 2 days) as compared to nonseeded control grafts. No grafts demonstrated aneurysm formation or dilatation in these short term implants.60 Compound Partially Bioresorbable Vascular Grafts Others have taken an intermediate approach and developed compound, or partially resorbable, vascular prostheses combining resorbable and permanent materials. Early studies were performed by Wesolowski’s group in the late 1950s and 1960s. They evaluated 24 different combinations of graft materials in dogs and pigs. They found, in general, that coated grafts performed less favorably than compound grafts wherein the resorbable component was an integral part of the fabrication. Deleterious changes (calcification) did occur in the compound grafts; however, this occurred in a delayed fashion thought to be a result of the resorbable component. The most satisfactory results were achieved with grafts composed of a compound yarn of resorbable catgut that had been wrapped with multi-filament polyester yarn.5 More recent work on compound vascular grafts has focused largely on the of incorporation of lactide polymers. An early report came from Ruderman’s group at Walter Reed. They evaluated woven grafts composed of 24% PLA and 76% Dacron implanted into the abdominal aorta of dogs. After 100 days they found all grafts to be patent and to show extensive tissue ingrowth.63 Compound Grafts: PG910 Bowald’s group investigated compound grafts as well. They compared double layer Vicryl grafts with Vicryl grafts anchored inside PTFE or Dacron mesh tube grafts. The grafts were implanted into the descending aorta of pigs. Nearly half of the pigs receiving Vicryl grafts experienced graft rupture. The pigs receiving grafts with external support did not experience graft rupture. Histologically, the supported Vicryl grafts were found to have the microscopic picture of arterial regeneration, although areas of degenerative change and calcification were noted. Increased collagen formation was noted in the Vicryl + Dacron group.16 Our group as well investigated Vicryl (PG910) in combination with Dacron and ePTFE.17,31 In addition, we investigated combinations of Vicryl and polypropylene. 64-66 Vicryl grafts surrounded by Dacron or PTFE showed no incidence of aneurysmal dilatation as compared to a 20% dilatation rate for unwrapped Vicryl grafts. Arterial regeneration occurred in the wrapped grafts, but in an incomplete manner. Grafts were also fabricated from compound yarns composed of blends of Vicryl and Dacron. The rate of aneurysmal dilatation increased proportionally with the percentage of Dacron in the blended yarn. A maximum dilation rate of 83% was seen with the 56% Dacron grafts. No

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dilatation was seen when 20% Dacron was used. Cellularity of the regenerated tissue in general was decreased in the presence of Dacron, suggesting an inhibitory effect of the biomaterial.17 Subsequent work demonstrated decreased prostaglandin content of the inner capsules of Dacron/Vicryl composite grafts as compared to those of Vicryl grafts.31 In order to avoid the deleterious effects of Dacron, compound grafts were constructed using polypropylene, which possesses greater biocompatibility and a incites only a relatively low-grade inflammatory response.65 Grafts fabricated from yarns containing 69% Vicryl (PG910) and 31% polypropylene were implanted into rabbit infrarenal aortas. Over 2 weeks to 12 months all grafts remained patent without aneurysm formation. Only residual polypropylene remained in the prostheses after 2 months. Production of 6-keto-PGF1α was in the normal range.64 PG910/polypropylene prostheses were also implanted into the aorto-iliac positions of dogs. Again, all grafts were without aneurysm formation over one year. Patency was enhanced as compared to Dacron and ePTFE control grafts. Arterial regeneration was demonstrated.65 Compound Grafts: PGA Our group also evaluated grafts made of polyglycolic acid (PGA) fabric reinforced by an outer wrap of Dacron. In this series, the incidence of aneurysmal dilatation was zero with or without the outer wrap up to nine months after implant. Again, Dacron appeared to have an inhibitory effect on the cellularity of the luminal surface and inner capsule.17 Evaluation of an advanced PGA based compound prosthesis is currently underway in our lab. Chu’s group at Cornell also investigated the combination of PGA and Dacron. They produced knitted fabric grafts composed of PGA and Dacron fibers blended at various compositional ratios. They studied the properties of these bicomponent fabrics in vitro. They felt that their most important finding was the achievement of increasing water porosity over time without significant losses in the structural integrity and strength of the specimens.67 Having demonstrated the effectiveness of partially resorbable arterial prostheses in rabbits, we sought to evaluate longer grafts in a higher order species. 4 mm by 5 cm conduits were woven from composite yarns containing 70% PDS/30% polypropylene and implanted into the aorto-iliac positions of dogs for periods of up to one year. No aneurysms or perigraft hematomas developed. The patency of the PDS/polypropylene grafts was significantly greater than that of Dacron or ePTFE control grafts. Inner capsules were completely endothelialized by 1 month. IC cellularity and thickness were greater than those within Dacron or ePTFE.65 Compound Grafts: PU/PLLA The Groningen group investigated compound grafts of PTFE fitted around PU/PLLA. They found that these grafts, as compared to grafts composed entirely of bioresorbable PU/PLLA or biostable but compliant PU, had diminished regeneration of all layers of the arterial wall, most strikingly in terms of elastic lamina. They attributed this

Tissue Engineering of Prosthetic Vascular Grafts

finding to the decreased compliance associated with the PTFE wrap.40 Compound Grafts: Polyethylene Oxide and Polylactic Acid Copolymers (PELA) The Biomaterials Research Laboratory at The Hebrew University of Jerusalem has developed and evaluated compound vascular prostheses fabricated from resorbable block copolymers of polyethylene oxide and polylactic acid (PELA) applied as a coating to knitted Dacron and polether urethane urea (Lycra) grafts. PELA is an elastomer that fully degrades in 3 weeks after implantation.68-71 The in vivo performance of the Dacron grafts was evaluated in dogs as right ventricle to pulmonary artery conduits. Unlike control woven grafts, the PELA coated grafts showed complete transmural tissue ingrowth after 14 months.68,69 In a separate study, the Hebrew University group compared ePTFE grafts to grafts comprised of Lycra fibers coated with PELA. 6 mm by 6 cm grafts were implanted in the canine carotid artery for 90 days. Unlike the ePTFE grafts, the PELA coated grafts remained compliant and demonstrated successful healing and incorporation. Further work with small caliber prostheses was anticipated.70

Resorbable Outer Wraps A novel application of resorbable materials in vascular grafting has been reported by Moritz and the Groningen group.72-75 Moritz applied a constrictive mesh tube of resorbable mesh to dilated, otherwise unusable veins with the goal of inducing neoarterial wall growth. He found that the meshes can effectively reduce the diameter of a venous graft by providing external support to the vessel wall, thus limiting distension.72 The Groningen group has also investigated biodegradable prostheses as external supports for vein grafts. They sought to prevent vein graft wall damage due to the higher arterial pressures encountered after implantation. They found that the prostheses can function as a protective scaffold for vein grafts in the arterial circulation, reducing damage to the vein graft wall and allowing gradual arterialization.73-75

Biopolymers: Resorbable Prostheses of Biologic Origin Another novel approach has been taken by several groups in Europe and the United States who have sought to develop resorbable prostheses through the chemical modification of arterial xenografts.76-81 Vascular prostheses of biological origin have been shown to develop aneurysms within several years of implantation. It has been hypothesized that, if properly processed, xenografts can function as a bioresorbable scaffold. Should this occur, an orderly regeneration of structural elements during healing could take place, the result being continued mechanical integrity of the graft.78 The biopolymer that has received the most attention is aldehyde crosslinked collagen.76 Moczar’s group in France developed a biodegradable microarterial graft from rat aorta. Trypsin treated arterial segments were coated with heparin or chondroitin sulfate to reduce thrombogenicity. Grafts

Bioresorbable Grafts: A Counterintuitive Approach

were implanted into the infrarenal aorta of rats for 3-8 weeks. The grafts were found to undergo resorption in parallel with re-endothelialization and scar tissue formation. Patency was around 50%.76 Macromolecular repair of the host aortic wall was documented in one year explants by the incorporation pattern of [3H]valine into proteins and the demonstration of de novo elastin synthesis in the scar replacing the prosthesis.77

Conclusion The resorbable graft represents a paradigm shift in vascular prostheses. Unlike conventional prostheses whose effectiveness depends by and large upon an inherent resistance to a natural healing process, resorbable grafts work with the body, providing a framework for well established healing patterns. It seems intuitive that by augmenting rather than opposing these patterns a more efficacious vascular graft should result. Great strides have been made toward this end. Further research will yield grafts for clinical evaluation in the near future. References 1. Shalaby S. Degradable materials: Perspectives, issues and opportunities. In: Barenberg S, Brash J, Narayan R, Redpath A, eds. The First International Scientific Consensus Workshop on Degradable Materials. Toronto: CRC Press, 1989:678. 2. Lommen E, Gogolewski S, Pennings AJ, Wildevuur CR, Nieuwenhuis P. Development of a neo-artery induced by a biodegradable polymeric vascular prosthesis. Trans Am Soc Artif Intern Organs 1983; 29:255-9. 3. Friedman S. A History of Vascular Surgery. Mount Kisco, New York: Futura Publishing Company, Inc., 1989. 4. Dennis C. Brief history of development of vascular grafts. In: Sawyer P, ed. Modern Vascular Grafts. New York: McGraw-Hill, 1987. 5. Wesolowski SA et al. The compound prosthetic vascular graft. A pathologic survey. Surgery 1963; 53:19. 6. Fries CC, Wesolowski SA. The polyester-oxidized cellulose compound vascular prostheses: A preliminary report. Trans Am Soc Artif Intern Organs 1964;X:227-30. 7. Robinson PH, van der Lei B, Knol KE, Pennings AJ. Patency and long-term biological fate of a two-ply biodegradable microarterial prosthesis in the rat. Br J Plast Surg 1989; 42:544-9. 8. Humphries A et al. Arterial prosthesis of collagen-impregnated dacron. Surgery 1961; 50:947. 9. Bascon J. Gelatin sealing to prevent blood loss from knitted arterial grafts. Surgery 1961; 50:504. 10. Guthrie C. End-results of arterial restitution with devitalized tissue. J Am Med Assoc 1919; 73:186. 11. Pachence J, Kohn J. Biodegradable polymers for tissue engineering. In: Lanza R, Langer R, Chick W, eds. Principles of Tissue Engineering. Georgetown, Texas: RG Landes Company, 1997. 12. Greisler HP. Bioresorbable materials and macrophage interactions. J Vasc Surg 1991; 13:748-50. 13. Bowald S, Busch C, Eriksson I. Arterial grafting with polyglactin mesh in pigs [letter]. Lancet. 1978; 1:153. 14. Bowald S, Busch C, Eriksson I. Arterial regeneration following polyglactin 910 suture mesh grafting. Surgery. 1979; 86:722-9.

501 15. Audell L, Bowald S, Busch C, Eriksson I. Polyglactin mesh grafting of the pig aorta. The two-year follow-up in an experimental animal. Acta Chir Scand 1980; 146:97-9. 16. Bowald S, Busch C, Eriksson I. Absorbable material in vascular prostheses: A new device. Acta Chir Scand 1980; 146:391-5. 17. Greisler HP, Schwarcz TH, Ellinger J, Kim DU. Dacron inhibition of arterial regenerative activities. J Vasc Surg 1986; 3:747-56. 18. Greisler HP, Dennis JW, Endean ED, Ellinger J, Buttle KF, Kim DU. Derivation of neointima in vascular grafts. Circulation. 1988; 78:I6-12. 19. Greisler HP. Arterial regeneration over absorbable prostheses. Arch Surg 1982; 117:1425-31. 20. Greisler HP, Kim DU, Price JB, Voorhees AB Jr. Arterial regenerative activity after prosthetic implantation. Arch Surg 1985; 120:315-23. 21. Leibovich SJ, Ross R. The role of the macrophage in wound repair. A study with hydrocortisone and antimacrophage serum. Am J Pathol 1975; 78:71-100. 22. Martin BM, Gimbrone MA Jr, Unanue ER, Cotran RS. Stimulation of nonlymphoid mesenchymal cell proliferation by a macrophage-derived growth factor. J Immunol 1981; 126:1510-5. 23. Greenburg GB, Hunt TK. The proliferative response in vitro of vascular endothelial and smooth muscle cells exposed to wound fluids and macrophages. J Cell Physiol 1978; 97:353-60. 24. Zarge J, Huang P, Greisler H. Blood vessels. In: Lanza R, Langer R, Chick W, eds. Principles of Tissue Engineering. Georgetown, Texas: RG Landes Company; 1997. 25. Greisler HP, Ellinger J, Schwarcz TH, Golan J, Raymond RM, Kim DU. Arterial regeneration over polydioxanone prostheses in the rabbit. Archives of Surgery 1987; 122:715-21. 26. Vorp DA, Raghavan ML, Borovetz HS, Greisler HP, Webster MW. Modeling the transmural stress distribution during healing of bioresorbable vascular prostheses. Ann Biomed Eng 1995; 23:178-88. 27. Greisler HP, Petsikas D, Lam TM et al. Kinetics of cell proliferation as a function of vascular graft material. J Biomed Mater Res 1993; 27:955-61. 28. Brant AM, Chmielewski JF, Hung T-K, Borovetz HS. Simulation in-vitro of pulsatile vascular hemodynamics using a CAD/CAM designed cam disk and roller follower. Artif Organs 1986; 10:419-421. 29. Pham S, S.J. D, R. J et al. Compliance changes in bioresorbable vascular prostheses following implantation. Surg Forum 1988; 39:330-332. 30. Greisler HP, Joyce KA, Kim DU, Pham SM, Berceli SA, Borovetz HS. Spatial and temporal changes in compliance following implantation of bioresorbable vascular grafts. J Biomed Mater Res 1992; 26:1449-61. 31. Schwarcz TH, Nussbaum ML, Ellinger J, Kim DU, Greisler HP. Prostaglandin content of tissue lining vascular prostheses. Curr Surg 1987; 44:18-21. 32. Greisler HP, Cabusao EB, Lam TM, Murchan PM, Ellinger J, Kim DU. Kinetics of collagen deposition within bioresorbable and nonresorbable vascular prostheses. ASAIO Trans 1991; 37:M472-5. 33. Greisler HP, Ellinger J, Henderson SC et al. The effects of an atherogenic diet on macrophage/biomaterial interactions. J Vasc Surg 1991; 14:10-23. 34. Lam TM, Whereat NE, Henderson SC, Burgess WH, Shaheen A, Greisler HP. Effects of hypercholesterolemia

502 on monokine-induced smooth muscle cell proliferation. Exs 1992; 61:346-56. 35. Greisler HP, Henderson SC, Lam TM. Basic fibroblast growth factor production in vitro by macrophages exposed to Dacron and polyglactin 910. J Biomater Sci Polym Ed. 1993; 4:415-30. 36. Zenni GC, Ellinger J, Lam TM, Greisler HP. Biomaterialinduced macrophage activation and monokine release. J Invest Surg 1994; 7:135-41. 37. Gogolewski S, Pennings A. Porous biomedical materials based on mixtures of polylactides and polyurethanes. Makromol Chem Rapid Commun 1982; 3:839-845. 38. Gogolewski S, Pennings AJ, Lommen E, Wildevuur CR. Small-caliber biodegradable vascular grafts from Groningen. Life Support Syst 1983; 1:382-5. 39. van der Lei B, Bartels HL, Nieuwenhuis P, Wildevuur CR. Microporous, complaint, biodegradable vascular grafts for the regeneration of the arterial wall in rat abdominal aorta. Surgery 1985; 98:955-63. 40. van der Lei B, Wildevuur CR, Nieuwenhuis P. Compliance and biodegradation of vascular grafts stimulate the regeneration of elastic laminae in neoarterial tissue: an experimental study in rats. Surgery 1986; 99:45-52. 41. van der Lei B, Darius H, Schror K, Nieuwenhuis P, Molenaar I, Wildevuur CR. Arterial wall regeneration in small-caliber vascular grafts in rats. Neoendothelial healing and prostacyclin production. J Thorac Cardiovasc Surg 1985; 90:378-86. 42. van der Lei B, Wildevuur CR, Nieuwenhuis P et al. Regeneration of the arterial wall in microporous, compliant, biodegradable vascular grafts after implantation into the rat abdominal aorta. Ultrastructural observations. Cell Tissue Res 1985; 242:569-78. 43. van der Lei B, Nieuwenhuis P, Molenaar I, Wildevuur CR. Long-term biologic fate of neoarteries regenerated in microporous, compliant, biodegradable, small-caliber vascular grafts in rats. Surgery 1987; 101:459-67. 44. van der Lei B, Wildevuur CR, Dijk F, Blaauw EH, Molenaar I, Nieuwenhuis P. Sequential studies of arterial wall regeneration in microporous, compliant, biodegradable small-caliber vascular grafts in rats. J Thorac Cardiovasc Surg. 1987; 93:695-707. 45. Hanson SJ, Jamshidi K, Eberhart RC. Mechanical evaluation of resorbable copolymers for end use as vascular grafts. ASAIO Trans 1988; 34:789-93. 46. Gogolewski SG, G. Degradable, microporous vascular prosthesis from segmented polyurethane. Colloid Polym Sci 1986; 264:854-8. 47. Greisler HP, Klosak JJ, Endean ED, McGurrin JF, Garfield JD, Kim DU. Effects of hypercholesterolemia on healing of vascular grafts. J Invest Surg 1991; 4:299-312. 48. Galletti G, Gogolewski S, Ussia G, Farruggia F. Long-term patency of regenerated neoaortic wall following the implant of a fully biodegradable polyurethane prosthesis: Experimental lipid diet model in pigs. Ann Vasc Surg 1989; 3:236-43. 49. Galletti G, Ussia G, Farruggia F, Baccarini E, Biagi G, Gogolewski S. Prevention of platelet aggregation by dietary polyunsaturated fatty acids in the biodegradable polyurethane vascular prosthesis: An experimental model in pigs. Ital J Surg Sci 1989; 19:121-30. 50. van der Lei B, Wildevuur CR. From a synthetic, microporous, compliant, biodegradable small-caliber vascular graft to a new artery. Thorac Cardiovasc Surg 1989; 37:337-47.

Tissue Engineering of Prosthetic Vascular Grafts 51. Greisler HP, Endean ED, Klosak JJ et al. Polyglactin 910/ polydioxanone bicomponent totally resorbable vascular prostheses. J Vasc Surg 1988; 7:697-705. 52. Leenslag JWK, Michel T, Pennings AJ, Van der Lei B.A complaint, biodegradable vascular graft: basic aspects of its construction and biological performance. New Polym Mater 1988; 1:111-26. 53. Pennings AJK, K.E., Hoppen HJ, Leenslag JW, Van der Lei B. A two-ply artificial blood vessel of polyurethane and poly(L-lactide). Colloid Polym Sci 1990; 268:2-11. 54. Galletti PM, Ip TK, Chiu TH, Nyilas E, Trudeli LA, Sasken H. Extending the functional life of bioresorbable yarns for vascular grafts. Trans Am Soc Artif Intern Organs 1984; 30:399-400. 55. Galletti PM, Trudell LA, Chiu TH et al. Coated bioresorbable mesh as vascular graft material. Trans Am Soc Artif Intern Organs 1985; 31:257-63. 56. Galletti PM, Aebischer P, Sasken HF, Goddard MB, Chiu TH. Experience with fully bioresorbable aortic grafts in the dog. Surgery 1988; 103:231-41. 57. Greisler HP, Kim DU. Vascular grafting in the management of thrombotic disorders. Semin Thromb Hemost 1989; 15:206-14. 58. Greisler HP, Klosak J, Dennis JW et al. Endothelial cell growth factor attachment to biomaterials. ASAIO Trans 1986; 32:346-9. 59. Wildevuur CR, van der Lei B, Schakenraad JM. Basic aspects of the regeneration of small-calibre neoarteries in biodegradable vascular grafts in rats. Biomaterials 1987; 8:418-22. 60. Yue X, van der Lei B, Schakenraad JM et al. Smooth muscle cell seeding in biodegradable grafts in rats: A new method to enhance the process of arterial wall regeneration. Surgery 1988; 103:206-12. 61. Greisler HP, Klosak JJ, Dennis JW, Karesh SM, Ellinger J, Kim DU. Biomaterial pretreatment with ECGF to augment endothelial cell proliferation. J Vasc Surg 1987; 5:393-9. 62. Gray JL, Kang SS, Zenni GC et al. FGF-1 affixation stimulates ePTFE endothelialization without intimal hyperplasia. J Surg Res 1994; 57:596-612. 63. Ruderman RJH, Andrew F, Hattler BG, Leonard F. Partially biodegradable vascular prosthesis. Trans Am Soc Artif Intern Organs 1972; 18:30-7. 64. Greisler HP, Kim DU, Dennis JW et al. Compound polyglactin 910/polypropylene small vessel prostheses. Journal of Vascular Surgery 1987; 5:572-83. 65. Greisler HP, Tattersall CW, Klosak JJ, Cabusao EA, Garfield JD, Kim DU. Partially bioresorbable vascular grafts in dogs. Surgery 1991; 110:645-54. 66. Zenni GC, Gray JL, Appelgren EO et al. Modulation of myofibroblast proliferation by vascular prosthesis biomechanics. Asaio J 1993; 39:M496-500. 67. Chu CCL, L.E. Design and in vitro testing of newly made bicomponent knitted fabrics for vascular surgery. Polym Sci Technol 1987; 35:185-213. 68. Cohn DY, Hani, Appelbaum Y, Uretzky G. A selectively biodegradable vascular graft. Prog Biomed Eng 1988; 5:73-9. 69. Uretzky G, Appelbaum Y, Younes H et al. Long-term evaluation of a new selectively biodegradable vascular graft coated with polyethylene oxide-polylactic acid for right ventricular conduit. An experimental study. J Thorac Cardiovasc Surg 1990; 100:769-76. 70. Cohn D, Elchai Z, Gershon B et al. Introducing a selectively biodegradable filament wound arterial prosthesis: A

Bioresorbable Grafts: A Counterintuitive Approach short-term implantation study. J Biomed Mater Res 1992; 26:1184-204. 71. Hellener G, Cohn D, Marom G. Elastic response of filament wound arterial prostheses under internal pressure. Biomaterials 1994; 15:1115-21. 72. Moritz A, Grabenwoger F, Windisch A et al. A method for constricting large veins for use in arterial vascular reconstruction. Artif Organs 1990; 14:394-8. 73. Zweep HP, Satoh S, van der Lei B et al. Autologous vein supported with a biodegradable prosthesis for arterial grafting. Ann Thorac Surg 1993; 55:427-33. 74. Zweep HP, Satoh S, van der Lei B, Hinrichs WL, Feijen J, Wildevuur CR. Degradation of a supporting prosthesis can optimize arterialization of autologous veins. Ann Thorac Surg 1993; 56:1117-22. 75. Hinrichs WL, Zweep HP, Satoh S, Feijen J, Wildevuur CR. Supporting, microporous, elastomeric, degradable prostheses to improve the arterialization of autologous vein grafts. Biomaterials 1994; 15:83-91.

503 76. Moczar M, Godeau G, Robert AM, Moczar E, Loisance D, Bessous JP. Biodegradable arterial prosthesis from rat aorta. Pathol Biol (Paris) 1980; 28:517-24. 77. Moczar M, Bessous JP, Loisance D. Healing of biodegradable vascular prostheses. Incorporation of 3H-Valine into proteins in the subendothelial scar and host intima-media of rat aorta. Connective Tissue Research 1983; 12:33-42. 78. Schmitz-Rixen T, Megerman J, Anderson JM et al. Longterm study of a compliant biological vascular graft. Eur J Vasc Surg 1991; 5:149-58. 79. Walter M, Erasmi H, Schmidt R. A new biological vascular prosthesis. Vasa Suppl 1991; 33:90-1. 80. Walter M, Erasmi H. A new vascular prosthesis of bovine origin. Thorac Cardiovasc Surgeon 1992; 40:38-41. 81. Walter M, Asoklis S, Erasmi H, Meisch JP. Bovine heterologous graft in replacement of a small calibre artery. Ann Chir 1994; 48:870-5.

Scaffold Engineering Material Aspects

CHAPTER 46 Biodegradable Materials K.J.L. Burg, S.W. Shalaby

Introduction

A

bsorbable materials are unique as implants, in that they are absorbed and excreted from the body at the conclusion of their functional period, thus alleviating the expense and potential complications of a retrieval surgery. Additionally, as these materials gradually absorb, they allow a gradual shift in mechanical strength from the biomaterial back to the healing tissue. The absorption time and therefore loss of mechanical strength of the implant may be modulated to best match a given application. Synthetic absorbable materials were first developed as medical implants in the late 1960s and found use in applications such as barriers, scaffolds, and drug delivery systems.1 Prior to this time, naturally derived materials such as collagen, catgut, and silk were used; however, they were subject to unpredictable, enzymatic degradation and elicited immune responses due to impurities. The evaluation and application of absorbable materials gradually expanded from soft tissue devices into bone fracture fixation devices, an area of ongoing development.2-5 A more recent area of interest is the role of absorbable materials as structural matrices or supports in tissue engineering.6,7

Criteria for Useful Tissue Engineering Materials Tissue engineering involves organ development and replacement by the seeding of cells on or into a polymer matrix which may be either cultivated in vitro or implanted directly in vivo. The selective use of parenchymal cells minimizes the amount of donor tissue required and, rather than implanting a large volume of avascular tissue which may not survive, the revascularization process occurs simultaneously with the tissue growth and material bioabsorption. Absorbable materials have found an important role as a matrix material in such constructs. Indeed, the advent of tissue engineering technology opened yet another area of absorbable materials research. Absorbable and biodegradable polymer technology lends itself well to tissue engineering in that it provides a means of creating a temporary, prefabricated tissue template into or on which cells may be seeded. The absorbable polymer construct is a temporary, structural support allowing and protecting new tissue growth and gradually absorbing as the tissue acquires the desired strengths and functions. Development of the tissue may be begun ex vivo,8 where the polymer supports and maintains the tissue integrity during implantation. The early attempts to grow tissue combined cells with collagenous two-dimensional substrates.9-11 Research has become diverse, with replacement potential in many areas including liver, skin, cartilage, breast, nerve, and bone repair. Along with the growth in possible applications came also the development of more suitable forms for three-dimensional Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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tissue culture. This included injectable gels and gel formers, fibrous scaffolds, porous foams, and channeled substrates to allow vascularization. The processing techniques also developed to allow more custom-made designs and provision of more complicated, site-specific shapes. The processing of an absorbable material allows a reproducible end product, and the use of an absorbable material reduces both the cost and the risk of donor and recipient surgery. There are many synthetic polymers which are used in tissue engineering research, among which are polylactide, polyglycolide, polyphosphazene, polycaprolactone, and various copolymers. Copolymers such as lactide-colysine have been specifically developed in order to provide attachment points for cell binding peptide groups.12-14 The synthetic polymers are advantageous in that their properties may be directly modulated during synthesis and processing to achieve particular requirements. They tend to degrade in a very predictable fashion, largely hydrolytically, which has obvious benefits. The natural materials which have found use in tissue engineering research include alginate, collagen, chitosan and coral, as well as similarly peptide modified natural materials (Table 46.1). These materials are naturally derived and therefore readily available; their absorption is generally enzymatically driven and subsequently less predictable. Typically this group of absorbable materials may vary considerably from batch to batch. They do, however, allow gelatinous, injectable structures and therefore a broader range of applications. There are many factors to consider when designing an absorbable substrate for tissue engineering purposes. The amount of exposed polymeric surface area will relate to the intensity of the inflammatory response; therefore, an implant with greater exposed surface will induce a greater response than a smaller area.7 The polymer purity, crystallinity, molecular orientation, molecular weight, and polydispersity may all influence the presence and timing of an inflammatory response. The effect of bulk mass loss of the polymeric material on the success rate of the newly developed tissue is, as a result, an important design issue. The pore size of a three-dimensional scaffold, as well as the means of dispersion and flow rate of cells into the scaffold, is of importance in determining both the distribution of cells and the integrity of the cells.15 Additionally, pore

sizes larger than 10 micron will allow invasion of fibrovascular tissue and therefore enhance capillary network development and subsequent vascularization.16 The rate of the tissue infiltration increases with pore size and porosity, and its survival and integration with the host tissue depends largely on the vascularization. An even distribution of cells throughout the matrix is important to the development and functionality of the tissue, as well as the interaction of the device and transplanted cells with the surrounding tissue. The hydrophilicity of the polymeric surface contributes to the wettability of the porous structure and hence to the distribution of cells throughout the matrix. The crystalline polylactide and polyglycolide materials are relatively hydrophobic; therefore, aqueous cell solutions do not adequately infiltrate the polymeric interior. Less polar solvents such as ethanol may be used to prewet the scaffold before applying the cell solution;17 the concern then becomes the minimization of the scaffold-solvent contact time in order to avoid excessive ester bond cleavage. Another approach to improve cellular distribution throughout the scaffold without addressing the bulk substrate properties is to surface coat the device with an absorbable, hydrophilic material prior to cell seeding. Low molecular weight, soluble polyvinyl alcohol coatings, as well as pluronic coatings, may be successfully applied to these hydrophobic substrates, thus rendering them more attractive for cellular attachment.18 These surfactants do not improve the cellular adhesion to the polymeric structure, but they do allow a much better cell distribution throughout the matrix. Furthermore, they dissolve readily in aqueous environment and are therefore unlikely to interfere with the intended implant function. The criteria for a tissue engineered absorbable vascular graft material are very focused. The material must be readily sculpted into an appropriately sized tubular shape, matching that of surrounding target tissue and, more problematic in the case of small diameter vessels, must be able to withstand large flow forces without hindering transport. The construct must be a specific porosity in order to optimize the tissue integration and reformation. The tissue ingrowth must occur over the entire length of the prosthesis rather than primarily from a transanastomotic pannus ingrowth. Most importantly, it must have an anti-thrombogenic lumenal cellular lining and must facilitate long term viability.

Table 46.1. Absorbable materials used in tissue engineering Absorbable Meterial

Material Type

Typical Forms

polylactide polyglyolide alginate collagen substituted polyphosphazene polycaprolactone poly(ethylene glycol-b-propylene glycol)

polyester polyester polysaccharide protein polyphosphazene polyester polyether

porous sponges fibrous meshes hydrogels, spheres porous sponges, knitted fabrics porous sponges porous sponges surfactant

Biodegradable Materials

Methods of Fabrication of Synthetic Polymeric Materials for Use in Tissue Engineering Tissue engineering absorbable graft research has thus far targeted the synthetic absorbable materials so it is useful to understand the various methods of processing these particular materials into tissue engineering scaffolds. Physical Aggregation Physical aggregation is the formation of a porous, interconnected network with no chemical interlocks between the polymer matrix components; rather, a physical adhesion is formed. The pores formed by aggregation are typically uniform and of low size distribution. This may be accomplished in microsphere systems by forming an agglomerate of particles by adding a biocompatible plasticizing agent such as triethylcitrate.19 Physical aggregation of this kind therefore allows an interconnected, biocompatible, mechanically stable, porous structure. A second method of physical aggregation is by sintering. Tassels and felts have found initial success as transplantation surfaces in liver and cartilage regeneration studies; however, these structures may lack the necessary mechanical integrity to maintain a specific shape. Fibrous scaffolds, woven or random arrays, have also found application in tissue engineering, but their dimensional stability and hence retention of shape is questionable. The fibers may be sintered together to provide a stable structure in a manner similar to the processing of the “self reinforced” bone pins;4 however, the high temperature required can cause disorientation of the molecular order and subsequent buckling of the material.20 The fibers may, rather, be coated with a supportive material which maintains the scaffold shape during heat treatment, but which may be later dissolved from the system.20 Chemical Aggregation Chemical aggregation involves forming a porous, interconnected network by crosslinking, or chemically bonding, the polymer matrix components. This may be accomplished with a difunctional agent such as gluteraldehyde.19 The drawbacks to this include reduced porosity due to interstitial “filler” material as well as the potential introduction of a toxic substance (e.g., traces of glutaraldehyde) into the biological system.

507

Gas Dispersion The pores formed by dispersion are typically uniform and of low size distribution. Gases may be introduced into a polymer melt or a liquid monomer, either directly or as a by-product of a chemical reaction. When cooling of the polymer or polymerization of the monomer occurs, the gas bubbles become entrapped in the material, usually resulting in a closed pore system. If the melt is extruded, this can actually result in an open pore system or a “blown” foam.21 These foams typically are a larger pore sized system. Alternatively, a gas may be introduced into a polymeric solid at high pressure. Then when the solid is transferred to atmospheric conditions, the gas will expand, forming a spongelike material (Fig. 46.1). This does not require use of solvents and may successfully be accomplished with absorbable polylactide/polyglycolide copolymers and a carbon dioxide environment.22 Solid and Liquid Dispersion Solid or liquid “fillers”, such as thermoset polymer spheres21 or water,23 may also be purposefully entrapped in the polymer matrix upon solidification. Subsequent removal of the solid or liquid results in an open pore system, the pore size distribution of which is generally larger than that of the gas dispersion process. Absorbable foams may be manufactured using this technique.18,24 The polymer can be mixed, for example, with a solvent and sodium chloride particles of a predetermined size (Fig. 46.2). The solvent is evaporated, leaving a polymer matrix with entrapped particles. The salt is leached from the system by soaking the construct in water, thus yielding a porous structure. The solid dispersion pore size is variable according to the size of the filler utilized. Supercritical Fluid Extraction Absorbable polymers, specifically polyglycolides, polylactides, and polyanhydrides, may be combined with supercritical fluids to process open-pore systems.25 The supercritical fluid is added to the polymer under pressure, the pressure is then dropped to remove the fluid and render the material spongelike (Fig.46.3). The percent porosity is dependent on the pressure differential between ambient and processing. One huge benefit of this technique is that no organic solvents are required. Fig. 46.1. Foam formation by gas dispersion.

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Fig. 46.2. Foam formation by solid dispersion.

Fig. 46.3. Phase diagram indicating pathway (1-2) of foam formation by supercritical fluid extraction.

Crystallization-Induced Microphase Separation This method of foam fabrication26,27 calls for the addition of a solid solvent with high melting temperature to the polymeric material. The mixture is then heated to form an isotropic liquid system. Subsequent cooling causes the crystallization-induced microphase separation and the now solidified solvent can be sublimed or removed with a liquid solvent. This process has been successfully applied to the formation of open pore absorbable constructs, specifically polydioxanone and poly(ε-caprolactone-co-glycolide) porous structures (Fig. 46.4).

Biodegradable Polymers for Vascular Grafts— Past, Present, and Future Many attempts have been made at engineering tubular vascular grafts by providing a tubular, porous, polymer support structure. The structure must be mechanically sound so as not to collapse or interfere with blood transport through the conduit. The polylactides and polyglycolides have shown favorable responses to such mechanical stresses.

These materials can be copolymerized in order to provide a range of degradation times from weeks to years. The relative lack of biocompatibility of traditional synthetic materials and the persistent dilemma of clinically unacceptable small diameter (< 6mm) vascular grafts instigated attempts to modify the substrates and provide a more compatible material. One such attempt is the surface modification of synthetic tubular materials or type I collagen with endothelial cells and smooth muscle cells to provide a more biologically “friendly” implant.28-30 Other investigators have tried coating 90/10 poly(glycolide-co-L-lactide) based grafts with absorbable materials such as a mixture of poly-DLlactide/poly-(2,3-butylene maleate) or poly-(2,3-butylene fumarate)/poly-DL-lactide.31-33 Fumarate leads to a more uniform and predictable absorption rate—there is no evidence of complications due to aneurysm or rupture. Alternatively, the porosity of a tubular implant may be adjusted to provide a diminishing array of pore sizes from the exterior to the interior of the implant, thus potentially allowing increased capillarization and improved biocompatibility.34

Biodegradable Materials

509

Fig. 46.4. Scanning electron micrograph of absorbable foam formed by crystallization-induced microphase separation.

Traditionally the ideal vascular graft material was assumed to be biologically inert. Today, however, research has shifted toward creating an active material that will interact and integrate favorably with the surrounding biological environment. The absorbable materials, as dynamic implant systems, have this quality and appear to stimulate growth factor release through macrophage induction. Both polyglycolide (PG), 90/10 glycolide/L-lactide copolymer, and polydioxanone (PDS) woven grafts have successfully demonstrated, in animal studies, inner capsule development as well as capillary invasion, thus resulting in an integrated, mechanically stable prosthesis.35-37 The inner capsule contains layers of myofibroblasts covered by endothelial-like cells (presumably from a capillary endothelial source). In similar studies, Dacron reinforced PGA shows no such capsule development and low amounts of transinterstitial macrophage migration, presumably due to the inhibitive nature of Dacron.38 These types of partially absorbable vascular patches, based on woven composite yarn comprising a 24-82% polyglycolide yarn, the balance being polyethylene terephthalate yarn, were also evaluated in dogs in terms of blood compatibility. 39 Results indicate that the patch biocompatibility and effect on blood components increased with the increase in the polyglycolide content. This signals the importance of tissue ingrowth to the construction of successful synthetic vascular grafts. Studies have also focused on the creation of an optimal biological component, integral to the prosthesis, for example, tubular nonabsorbable polyester fabrics incorporated with bovine serum albumin and a luminal coating of gelatin heparin sodium solution. These constructs are crosslinked with a polyepoxy compound. The grafts, once implanted, show no signs of platelet aggregation or fibrin formation and show development of neointima over the anastomotic graft site. The degree of crosslinking is critical

to success, as too high a degree causes lack of anastomotic endothelialization and transinterstitial tissue ingrowth.40 The compliance and longevity of a potential vascular graft material can influence the retention of mechanical integrity. 41,42 Studies in the polyurethane/polylactide (PU/PLLA) system have shown that a decrease in the percent polyurethane causes a decrease in compliance and an increase in aneurysmal degeneration.43 These studies further showed that high mechanical compliance of the graft at implantation stimulates elastin formation and an absorbable graft material allows maintenance of the mechanical stability of the newly formed tissue. The PU/PLLA system is initially a compliant one; however, long term incidences of aneurysmal dilation are significantly greater in this system than the polyester ones. This is possibly because the ongoing presence of PU leads to fibroplasia, resulting in decreased compliance and elastin deposition. Another important consideration for scaffold development, as with other applications, is that the tissues should gain adequate strength before the material absorbs. If this is not possible, the absorbable scaffold may be reinforced by an inert (not Dacron) component or developed into a composite material whereby the tissue ingrowth occurs through a series (two or more) of polymers. The multicomponent absorbable vascular prostheses composed of compound yarns of this type appear to be efficacious.44 Studies have shown that, in contrast to the absorbable component, the biomechanical properties of the nonabsorbable component modulate the histologic attributes and longevity of the final tissue structure. Other attempts at tissue engineering blood vessels have been made by constructing the vessels ex vivo directly from the cellular components. This may be accomplished by culturing a sheet of human vascular smooth muscle cells in collagen and placing this about a lumenal support to produce the media.45 A fibroblast sheet is then similarly cultivated

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Tissue Engineering of Prosthetic Vascular Grafts

Fig. 46.5. Surface phosphorylation of polyethylene.

and placed about the media to form the adventitia. Endothelial cells are later seeded into the lumen of the vessel, thus forming a mechanically sound three-dimensional vessel.

Newly Tailored Materials for Tissue Engineering Two recently developed technologies are expected to be of value in tissue engineering applications. First, Shalaby46 developed a series of injectable, absorbable liquid copolymers which undergo gel formation upon contacting moist biological tissues. These copolymers are based on polyethylene glycol segments and copolymeric segments derived from glycolide, lactide, trimethylene carbonate, ε-caprolactone, and/or p-dioxanone. These monomers are used in the production of commercial absorbable sutures. The injectable liquid polymers were described as a suitable vehicle for housing living cells for injection into desired biological sites to continue propagation into three-dimensional constructs with the gradual absorption of the gel carrier. Surface phosphonylation of preformed devices, including microporous open-cell foams that were described by Shalaby and co-workers,27,47,48 represents a second area of technology with promising potential for use in tissue engineering (Fig 46.5). This technology entails the formation of covalently bonded surface phosphonate groups which can: 1. Introduce hydrophilicity to hydrophobic substrates; 2. Bind specific proteins or peptides to direct cell attachment; and 3. Bind Ca2+ for directing the deposition of hydroxyapatite toward bone formation. Among the substrates suitable for surface phosphonylation are certain hydrophobic absorbable polyesters.

Future Prospectives One of the drawbacks to the synthetic materials as potential vascular graft materials is that, as homopolymers, they do not have the cell surface receptors required for cellular recognition. The polymers must be optimized to best modulate features such as crystallinity, surface texturing, and surface wettability in order to selectively attract specific biological components while preventing spreading of blood platelets. Surface modification such as that described by Shalaby and co-workers may be of value in overcoming some of these issues.

The need for a vascular graft implies that the surrounding target tissue may be “abnormal” (atherosclerotic, for example). It has been shown in animal studies that the quality of such tissues (for example, in the presence of excessive amounts of lipoprotein) may have a profound effect on the development of new tissue.49 It would therefore be useful to study the absorbable material in such suboptimal conditions or in compromised tissue environments. References 1. Gilding DK. Biodegradable polymers. In: Williams DF, ed. Biocompatibility of Clinical Implant Materials. Boca Raton, FL: CRC Press, 1981:209-232. 2. Bos RRM, Rozema FR, Boering G et al. In vivo and in vitro degradation of poly(l-lactide) used for fracture fixation. In: Putter C de, Lange GL de, Groot K de, Lee AJC, eds. Advances in Biomaterials. 8. Amsterdam, The Netherlands: Elsevier Science, 1988:245-250. 3. Leenslag JW, Pennings AJ, Bos RRM et al. Resorbable materials of poly(L-lactide). VI. Plates and screws for internal fracture fixation. Biomater 1987; 8(1):70-73. 4. Vainionpää S, Kilpikari J, Laiho J et al. Strength and strength retention in vitro, of absorbable, self-reinforced polyglycolide (PGA) rods for fracture fixation. Biomater 1987; 8(1):46-48. 5. Vert M, Christel P, Garreau H et al. Totally bioresorbable composite systems for internal fixation of bone fractures. In: Chiellini E, Giusti P, Migliaresi C, Nicolais L, eds. Polymers in Medicine II. New York, New York: Plenum Press, 1986:263-275. 6. Vacanti JP. Beyond transplantation. Arch Surg 1988; 123:545-549. 7. Vacanti JP, Morse MA, Saltzman WM et al. Selective cell transplantation using bioabsorbable artificial polymers as matrices. J Pedia Surg 1988; 23(1):3-9. 8. Mikos AG, Sarakinos G, Lyman MD et al. Prevascularization of porous biodegradable polymers. Biotechnol Bioeng 1993; 42:716-723. 9. Boyce ST, Christianson DJ, Hansbrough JF. Structure of a collagen-GAG dermal skin substitute optimized for cultured human epidermal keratinocytes. J Biomed Mater Res 1988; 22:939-957. 10. Boyce ST, Hansbrough JF. Biologic attachment, growth, and differentation of cultured human epidermal keratinocytes on a graftable collagen and chondroitin-6sulphate substrate. Surgery 1988; 103(4):421-431. 11. Hansbrough JF, Boyce ST, Cooper ML et al. Burn wound closure with cultured autologous keratinocytes and fibroblasts attached to a collagen-glycosaminoglycan substrate. JAMA 1989; 262(15):2125-2130.

Biodegradable Materials 12. Cook AD, Hrkach JS, Gao NN et al. Characterization and development of RGD-peptide-modified poly(lactic acidco-lysine) as an interactive, resorbable biomaterial. J Biomed Mater Res 1997; 35:513-523. 13. Barrera DA, Zylstra E, Lansbury PT et al. Copolymerization and degradation of poly(lactic acid-co-lysine). Macromolecules 1995; 28:425-432. 14. Hrkach JS, Ou J, Lotan N et al. Synthesis of poly(L-lactic acid-co-L-lysine) graft copolymers. Macromolecules 1995; 28:4736-4739. 15. Wald HL, Sarakinos G, Lyman MD et al. Cell seeding in porous transplantation devices. Biomater 1993; 14(4):270278. 16. Mooney DJ, Langer R. Engineering biomaterials for tissue engineering: The 10-100 micron size scale. In Bronzino JD, ed. Biomedical Engineering Handbook. Boca Raton, FL: CRC Press, 1995:1609-1618. 17. Mikos AG, Lyman MD, Freed LE et al. Wetting of poly(Llactic acid) and poly(DL-lactic-co-glycolic acid) foams for tissue culture. Biomater 1994; 15(1):55-57. 18. Mooney DJ, Park S, Kaufmann PM et al. Biodegradable sponges for hepatocyte transplantation. J Biomed Mater Res 1995; 29:959-965. 19. Schugens Ch, Grandfils Ch, Jerome R et al. Preparation of a macroporous biodegradable polylactide implant for neuronal transplantation. J Biomed Mater Res 1995; 29:1349-1362. 20. Mikos AG, Bao Y, Cima LG et al. Preparation of poly(glycolic acid) bonded fiber structures for cell attachment and transplantation. J Biomed Mater Res 1993; 27:183-189. 21. Frisch KC, Saunders JH, eds. Plastic Foams. New York: Marcel Dekker, 1972. 22. Mooney DJ, Baldwin DF, Suh NP et al. Novel approach to fabricate porous sponges of poly(DL-lactic-co-glycolic acid) without the use of organic solvents. Biomater 1996; 17(14):1417-1422. 23. Williams JM, Wrobleski DA. Microstructures and properties of some microcellular foams. J Mater Sci Letters 1989; 24:4062-4067. 24. Mikos AG, Thorsen AJ, Czerwonka LA et al. Preparation and characterization of poly(L-lactic acid) foams. Polymer 1994; 35:1068-1077. 25. De Ponti R, Torricelli C, Martini A et al. Use of supercritical fluids to obtain porous sponges of biodegradable polymers. International Patent WO 91/09079, June 1991. 26. Roweton SL. A new approach to the formation of tailored microcellular foams and microtextrued surfaces of absorbable and non-absorbable thermoplastic biomaterials. Master of Science Thesis. Department of Bioengineering, Clemson University. 1994. 27. Shalaby SW, Roweton SL. Continuous open-cell polymeric foams containing living cells. 1997: U.S. Patent #5,677,355. 28. Herring MB, Gardner AK, Glover JL. Seeding human arterial prostheses with mechanically derived endothelium. The detrimental effect of smoking. J Vasc Surg 1984; 1:279-289. 29. Yue X, van der Lei B, Schakenraad JM et al. Smooth muscle cell seeding in biodegradable grafts in rats: A new method to enhance the process of arterial wall regeneration. Surgery 1988; 103:206-212. 30. Kempczinski RF, Rosenman JE, Pearce WH et al. Endothelial cell seeding of a new PTFE vascular prosthesis. J Vasc Surg 1985; 2(3):424-429.

511 31. Galletti PM, Ip TK, Chiu T-H et al. Extending the functional life of bioresorbable yarns for vascular grafts. Trans ASAIO 1984; 30:399-400. 32. Galletti PM, Trudell LA, Chiu T-H et al. Coated bioresorbable mesh as vascular graft material. Trans ASAIO 1985; 31:257-263. 33. Galletti PM, Aebischer P, Sasken HF et al. Experience with fully bioresorbable aortic grafts in the dog. Surgery 1988; 103:231-241. 34. Lei Bvd, Blaau EH, Dijk F et al. Microporous compliant biodegradable graft materials: A new concept for microvascular surgery. In: Skotnicki SH, Buskens FGM, and Reinaerts HHM, eds. Recent Advances in Vascular Grafting. The Netherlands: Nijmegan, 1984:19-23. 35. Greisler HP. Arterial regeneration over absorbable prostheses. Arch Surg 1982; 177:1425-1431. 36. Greisler HP, Kim DU, Price JB et al. Arterial regenerative activity after prosthetic implantation. Arch Surg 1985; 120:315-323. 37. Greisler HP, Ellinger J, Schwarcz TH et al. Arterial regeneration over polydioxanone prostheses in the rabbit. Arch Surg 1987; 122:715-721. 38. Greisler HP, Schwarcz TH, Ellinger J et al. Dacron inhibition of arterial regenerative activity. J Vasc Surg 1986; 3:747-756. 39. Yu TJ, Ho DM, and Chu CC. Bicomponent vascular grafts consisting of synthetic absorbable fibers: Part II: in vivo healing response. J. Investig. Surg 1994; 7(3):195-211. 40. Niu S, Kurumatani H, Satoh S et al. Small vascular prostheses with incorporated bioabsorbable matrices. A preliminary study. ASAIO J 1993; 39(3):M750-M753. 41. Greisler HP, Joyce KA, Kim DU et al. Spatial and temporal changes in compliance following implantation of bioresorbable vascular grafts. J Biomed Mater Res 1992; 26(11):1449-61. 42. Greisler HP, Pham SM, Endean ED et al. Relationship between changes in biomechanical properties and cellular ingrowth in absorbable vascular prostheses. ASAIO Abstracts 1987; 16:25. 43. Lei Bvd, Bartels HL, Nieuwenhuis P et al. Microporous, compliant, biodegradable vascular grafts for the regeneration of the arterial wall in rat abdominal aorta. Surgery 1985; 98:955-963. 44. Endean ED, Kim DU, Ellinger J et al. Effects of polypropylene’s mechanical properties on histological and functional reactions to polyglactin 910/polypropylene vascular prostheses. Surg Forum 1987; 38:323-325. 45. Germain L, L’Heureux N, Labbé R et al. Human blood vessel produced in vitro by tissue engineering. Workshop on Biomaterials and Tissue Engineering Abstracts 1997; 23. 46. Shalaby SW. Hydrogel-forming self-solvating absorbable polyester copolymers and methods for use thereof. U.S. Patent 5,612,052 Assignee: Poly-Med, Inc. 1997. 47. Shalaby SW, McCaig SM. Process for phosphonylating the surface of an organic polymeric preform, U.S. Patent 5,491,198, Assignee: Clemson University 1996. 48. Shalaby SW, Rogers KR. Polymeric prosthesis having a phosphonylated surface, U.S. Patent 5,558,517, Assignee: Clemson University 1996. 49. Greisler HP, Klosak JJ, Endean ED et al. Effects of hypercholesterolemia on healing of vascular grafts. J Investigative Surg 1991; 4(3):299-312.

Scaffold Engineering Material Aspects

CHAPTER 47 The Influence of Porosity and Surface Roughness on Biocompatibility J.M. Schakenraad, K.H. Lam

Introduction

V

ascular prostheses might solve a lot of clinical problems, if only they would “do their job” in a functional way. Functional in this respect includes: nonthrombogenicity; compliance, diameter and wall thickness similar to vessel wall; blood compatibility in all its aspects (complement activation, red and white blood cell damage, thrombocyte activation etc.); immunological inertness etc. Some of these parameters can be influenced in such a way that an acceptable functionality is achieved. Especially with regard to small diameter vascular grafts (< 3 mm), it has proven impossible to solve all problems simultaneously. A vast challenge still lies ahead for the next generation of scientists. This chapter will discuss two material surface parameters (roughness and porosity) having a large influence on biocompatibility and ultimate patency of vascular grafts. A parameter indicative for short term biocompatibility is the inflammatory response towards an implanted biomaterial. Upon implantation of a biomaterial, local1,2 and systemic3 effects can be observed. The local tissue reaction consists of an inflammatory response which serves to eliminate the cause of an injury, minimize the damage and trigger mechanisms for repairing the tissue damaged by injury. The surgical procedure itself initially determines the type and intensity of the inflammatory response. However, in addition, implanted biomaterials themselves provoke an inflammatory response. Many material characteristics may influence the inflammatory response against a biomaterial (being either a natural polymer, a chemical polymer, metals, ceramics etc.): chemical composition,1 material toxicity4,5 surface free energy (wettability),6 surface charge7 surface morphology such as porosity,8 roughness of the surface9 and shape,7 implantation site,10 rate of degradation,11 degradation products,12,13 and many more. Surface morphology is one of the first factors directly influencing the tissue/inflammatory response. This is indicated in various studies.7,14 Matlaga et al demonstrated the role of shape.7 The intensity of the inflammatory response increases when the number of edges of the implanted materials increases. Also, there are studies indicating that the inflammatory response against implanted porous polymers or biomaterials is more intense when compared to nonporous materials.8 Wettability may influence the tissue reaction to the polymer, because there is a range in wettability values that is optimal for cell adhesion, growth and spreading. Cells attach and proliferate less well on polymers having a wettability which is too low15 or too high.16 Another factor is degradability. Degrading polymers provoke a more intense inflammatory response compared to nondegrading polymers. A possible cause for this observation Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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may be the release of monomers, oligomers and/or fragments upon degradation.12,13 However, at earlier stages of the degradation process, changes (such as increase) in the surface morphology of a polymer (film) may occur, which in turn may alter the inflammatory response. To illustrate the effect of surface morphology on the tissue response and ultimate biocompatibility of a material after implantation, the following series of materials are compared: 1. Porous versus nonporous; 2. Rough versus smooth; 3. Degradable versus nondegradable. The inflammatory response was characterized using semi-quantitative techniques comprising morphological criteria and monoclonal antibodies directed against epitopes which are specific for the respective cell types. Such aspects as molecular weight of soluble leachables and relative surface area are considered to be closely related to the observed histological parameters and are therefore also included in this overview.

Material Characterization Except for PTFE, the final thickness of the films was determined with scanning electron microscopy (SEM). All films were cleaned by washing in a phosphate buffered saline (PBS) for 24 h prior to use.

Evaluation Models Materials As a degradable material, the well known and widely used polylactic acid is used, both as a solid, a porous and a combination form. As a nondegradable reference material, PTFE is used, both in a solid shape and as the expanded PTFE with a pore size of 30 micrometers. Poly(L-lactic acid) (PLLA) films were cast from PLLA with a reported Mn of 50,000 (Purac Biochem B.V., The Netherlands).17 Three types of films were cast: a nonporous type, a porous type and a “combi” type (porous with a nonporous layer on one side). The base parameters of the PLLA films are listed in Table 47.1. All PLLA films were cut in strips of 15 x 2 mm. Polytetrafluoroethylene (PTFE) was obtained commercially (Wientjes, The Netherlands). PTFE was cut into strips of 15 x 2 x 1 mm. Expanded polytetrafluoroenthylene (ePTFE) was obtained as nonsterile GORE-TEX® ePTFE cell collector tubing (WL Gore & Associates GMBH, Germany). The tubing was cut open along the longitudinal axis to obtain films measuring 15 x 2 x 0.25 mm.

➔ Table 47.1. Base parameters of the PLLA films Parameter

Non-porous PLLA film

Porous PLLA film

Combi PLLA film

Mw Mn Mw/Mn Tm Heat of fusion

98,000 42,000 2.3 176°C 53 J/g

109,000 41,000 2.7 180°C 51 J/g

167,000 53,000 3.1 181°C 50 J/g

Fig. 47.1. Scanning electron micrographs of a cross-section of PLLA films. (A, top.) nonporous, (B, middle.) porous, (C, bottom.) “combi”; arrowhead indicates the nonporous side. The thickness of the films is: nonporous, 33 µm; porous, 244 µm and “combi”, 82 µm. The nonporous layer of the “combi” film was approximately 5 µm. The pore size of the porous film varied from approximately 1-150 µm and of the “combi” film from 1-50 µm.

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Surface Morphology The surface morphology of the different specimens is illustrated in Figure 47.1A-C. Specimens of all films were sputter-coated with gold (Balzers 07 120B) and their surface morphology was examined with DS 130 scanning electron microscope (SEM) (ISI), operated at 10 KV. The thickness of the polymers was calculated using SEM: nonporous 33 µm (Fig.47.1A); porous 244 µm (Fig. 47.1B); and the “combi” film 82 µm (Fig. 47.1C). The nonporous layer of the “combi” film was approximately 5 µm. The pore size of the porous film varied from approximately 1 to 150 µm and of the “combi” film from 1 to 50 µm. PTFE was nonporous, having the same morphological appearance of the surface as the nonporous PLLA film. The thickness of the porous ePTFE film was approximately 250 µm with a fibril length of approximately 90 µm (Fig. 47.2). Wettability Wettability, as a measure for surface free energy (hydrophobicity) was determined by contact angle measurements using the sessile drop technique described by Busscher et al.18 For PTFE and the nonporous PLLA, the contact angles were determined using H20 and α−bromonaphthalene as wetting agents. Five measurements were made for each polymer and wetting agent. The highest value of each set of five measurements is maximally 5° higher than the lowest value. Therefore, it was not relevant to determine the standard deviation. The H2O contact angles were 101° for PTFE and 72° for nonporous PLLA, and the α-bromonaphthalene contact angles were 66° for PTFE and 23° for nonporous PLLA. This shows that nonporous PLLA has a less hydrophobic surface than PTFE. Implantation Procedure The films were disinfected prior to implantation and implanted subcutaneously in 21 female (AO x BN)F1 rats obtained from our own breeding facility. In each rat a non-

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porous PLLA film, a porous PLLA film, a “combi” PLLA film, a PTFE film and an ePTFE film were implanted. The sixth incision and subcutaneous pocket served as a control (sham operation). Three samples per polymer and time interval were implanted. The rats had free access to standard rat food and water. All national rules concerning the care and use of laboratory animals have been observed. The rats were sacrificed after 1, 3, 7, 14, 40, 90 or 180 days and the polymer films were removed with excess surrounding tissue. Evaluation of the Inflammatory Response Light Microscopy After harvesting, the samples were immediately fixed for at least 24 h at 4°C in a 0.1 M Na-cacodylate buffer, pH 7.4, containing 2% glutaraldehyde and 0.1 M sucrose. The samples were then dehydrated in a graded ethanol series and embedded in glycolmethacrylate (Technovit®, Kulzer, Germany), allowing sections to be cut perpendicular to the longitudinal axis of the polymer film. Sections for light microscopy examination were cut on a microtome (Jung autocut 1140, equipped with a D knife with a tungsten carbide cutting edge), mounted on glass slides and stained with toluidine blue and alkaline fuchsin.19 Immunohistochemical Staining After harvesting, samples were snap-frozen at -80°C using liquid freon. Cryostat sections of 7 µm were cut, mounted on glass slides, air-dried and fixed in acetone for 12 minutes. The sections were then again air-dried for 1 h and incubated with the first-stage, cell type specific, monoclonal antibody (mAb) for 1 h. Subsequently, the sections were washed 3 times in PBS, followed by incubation with the second-stage antibody conjugated to peroxidase, diluted 1:40 in PBS and supplemented with 5% v/v normal rat serum to prevent nonspecific binding. Swine anti-rabbit Ig

Fig. 47.2. Scanning electron micrograph of (porous) ePTFE. The fibril length (distance between A and B) is approximately 90 µm. Bar indicates 30 µm.

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(Dakopatts, Denmark) was used as second-stage antibody to detect the first-stage mAb α-Asialo GM1 (Table 47.2). Rabbit anti-mouse Ig (Dakopatts, Denmark) was used to detect the other first-stage mAbs. After incubation with the second-stage antibody, sections were rinsed 3 times in PBS for 5 minutes. Peroxidase activity was demonstrated by applying 3,3-diaminobenzidine tetrahydrocloride (Sigma) at a concentration of 0.5 mg/ml in 0.05 M Tris-HCI buffer (pH 7.6) containing 0.01% H2O2 for 10 minutes. After rinsing in fresh tap water, sections were counterstained lightly with hematoxylin for 10 seconds. The sections were then dehydrated using a graded ethanol series and xylene and subsequently covered with coverslips using DePeX (Gurr, BDH Ltd, England) mounting medium. In controls, PBS was used instead of the first-stage mAb. The first-stage mAbs used, and their sources, are shown in Table 47.2. Quantification of the Inflammatory Response The magnification of the light microscope was set at x400 when examining the section stained using mAbs. The staining patterns of the tissue surrounding or invading the polymer film was evaluated and the number of positive cells surrounding or invading the polymer films per field of view counted. Four fields of view per section, two sections per sample and three samples for each period of implantation, polymer film and monoclonal antibody respectively were examined. The tissue reactions at the edges of the polymer films were excluded from evaluation, to avoid artifacts due to mechanical irritation. The number of cells staining positive per field of view was classified as follows: grade 0 = no positive cells, 1 = 1 to 5 positive cells per field of view, 2 = 5 to 10 positive cells per field of view, 3 = 10 to 25 positive cells per field of view and 4 = more than 25 positive cells per field of view. The average class of the 24 fields of view for each period of implantation, polymer film and monoclonal antibody respectively is reported.

Macroscopically, there were no signs of an infection demonstrated in any of the rats. The results of the semiquantitative evaluation of the immunohistochemical staining to detect leukocytes (OX 1), neutrophilic granulocytes (HIS 48), macrophages (ED 1, ED 2, ED 3), T lymphocytes (OX 19), vast majority of B lymphocytes (HIS 40), natural killer cells (α-Asialo GM1), cells expressing MHC class II antigen—such as activated fibroblasts—(HIS 19) are demonstrated in Figures 47.3A-I respectively. False positive cells were observed occasionally in PBS controls. This was due to endogenous peroxidase activity expressed only by neutrophilic granulocytes, since the staining pattern using HIS 48 corresponds with the pattern of endogenous peroxidase activity. Moreover, cells expressing endogenous peroxidase activity can be well distinguished due to their more intense staining as compared to cells stained for mAbs. The concentration of B lymphocytes (Fig. 47.3D) and the concentration of natural killer cells (Fig. 47.3H) was to a large extent similar and low for both PLLA, PTFE films and the subcutaneous tissue under the scar of the sham operation. At day3, the concentration of each cell type involved in the inflammatory response as a measure for the intensity was approximately the same for all the polymer films (Fig. 47.3A to 47.3G). At day 7, macrophages are beginning to play a prominent role in the inflammatory response (Fig. 47.3E). Subsets of macrophages surrounding the polymer films were found in different locations. ED 1 positive macrophages are found in the entire area, demonstrating inflammatory response against the implant, especially in the area in closest approximation to the polymer film (Fig. 47.4A). In contrast, ED 2 and ED 3 positive macrophages are neither found in the area in closest approximation the polymer film, nor in the pores of the PLLA or PTFE film. ED 2 and ED 3 positive macrophages remain restricted to the surrounding tissue at some distance from the implant (Fig. 47.4B and 47.4C). The relative ratio between the subsets was fairly constant for the respective implants.

Table 47.2. Antibodies (mAbs) used to examine and quantify the inflammatory response against the implanted polymer films Monoclonal Antibody

Epitope

Mainly characterizing cell type in subcutaneous tissue:

OX 1 HIS 48 OX 19 HIS 40 ED 1 ED 2 ED 3

CD 45 probably surface CD 5 IgM heavy chain Lysosomal antigen Surface antigen Surface antigen

α-Asialo GM1

Probably surface antigen MHC class II

All leukocytes Granulocytes T-lymphocytes B-lymphocyte subset likely to react first upon inflammatory response Majority of macrophages. Probably associated with active phagocytosis. Subset macrophages. Probably associated with maturity. Subset macrophages. Probably associated with downregulation of inflammatory reaction Large granular lymphocytes natural killer cells Activated tissue cells (fibroblast) IDC, subset macrophage

HIS 19

Source of the mAbs: OX antibodies were a generous gift of the late Dr. A.F. Williams, Department of Biochemistry, University of Oxford, Great Britain; ED antibodies were a generous gift of Dr. C.D. Dijkstra, Department of Cell Biology. The HIS and a-Asialo antibodies were generous gifts of Dr. P. Nieuwenhiu’s of the Department of Histology, University of Groningen.

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Fig. 47.3. The average class of 24 fields of view of the number of cells involved in the inflammatory response against the implanted polymer films, stained with different mAbs. Class criteria: grade 0 = no positive cells; grade 1 = 1 to 5 positive cells; grade 2 = 5 to 10 positive cells; grade 3 = 10 to 25 positive cells and grade 4 = more than 25 positive cells per field of view at an original magnification of x400. The following mAbs were used: (A, top.) OX 1, all leukocytes. (B, bottom.) HIS 48, neutrophilic granulocytes.

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Fig. 47.3. (C, top.) OX 19, T-lymphocytes, (D, bottom.) HIS 40, vast majority of B lymphocytes.

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Fig. 47.3. (E, top.) ED 1, vast majority of macrophages, including multinuclear giant cells; (F, bottom.) ED 2, subset (mature/resident) macrophages.

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Fig. 47.3. (G, top.) ED 3, subset macrophages. (H, bottom.) α -Asialo GM1, natural killer cells/large granular lymphocytes

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Fig. 47.3. (I.) HIS 19, cells expressing MHC II antigen (e.g., activated fibroblasts, macrophages, dendritic cells). X-axis represents implantation time (days), Y-axis the average number of cells per field of view as class 0 to 4. Z-axis represents the different polymer films.

In GMA sections it is observed that the inflammatory response at day 7 is becoming more intense (Fig. 47.5) for the porous PLLA and porous side of the “combi” PLLA films. For the nonporous PLLA and the nonporous side of the “combi” PLLA film, the onset of encapsulation by approximately 3 layers of fibroblasts was observed. Only a minimal encapsulation is observed for the porous PLLA film at day 7. In contrast to the porous PLLA film almost no cellular invasion of the ePTFE film was observed. The cell layer surrounding the ePTFE consists mainly of macrophages. The PTFE film was surrounded by one or two layers of macrophages and the onset of encapsulation can be observed. At day 14, the inflammatory response becomes chronic with predominantly macrophages and T lymphocytes surrounding the films. In the GMA sections, foreign body giant cells can be observed surrounding the PLLA fims (Fig. 47.6). All films are now encapsulated by continuous layers of fibrocytes and collagen. The fibrocytes are not stained by HIS 19 monoclonal antibody, indicating a decrease in cell activity. The porous PLLA film and porous side of the “combi” film provoke a more intense inflammatory response than the nonporous PLLA film and nonporous side of the “combi” film respectively. However, this observation could not be made for the ePTFE film as compared to the PTFE film. At day 40, in general, the intensity of the inflammatory response against the polymer films had decreased further. However, PLLA films still provoked a more intense in-

flammatory response than PTFE films as demonstrated by the higher concentration of neutrophils, macrophages/giant cells and T lymphocytes surrounding the PLLA films. The concentration of cells against PTFE and ePTFE films is comparable to the concentration of cells in the subcutaneous tissue under the scar of the sham operation. The inflammatory response against the nonporous PLLA film was mainly localized at the edges of the film or at the edges of the pieces when broken. In contrast, the inflammatory response against porous PLLA films was localized in the pores. The “combi” film shows a more pronounced inflammatory response at the porous side. At day 90, the difference in the intensity of the inflammatory response between PLLA and PTFE films and also between nonporous and porous PLLA films had become much more pronounced. This is demonstrated by the increased concentration of macrophages, leukocytes and cells expressing MHC class II (HIS 19) antigen, (probably activated fibroblasts) surrounding or invading the PLLA films. The porous PLLA films provoke a more intense inflammatory response than the nonporous PLLA film. In contrast, the inflammatory response against PTFE and ePTFE films remained the same as at day 40. At day180, the difference in the intensity of the inflammatory response between the porous and the nonporous PLLA film was more pronounced. In contrast, the tissue reaction (“inflammatory response”) against PTFE and ePTFE did not differ much from the tissue reaction in the

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Fig. 47.4. Immunohistochemical staining of the tissue surrounding the porous PLLA film (P) at day 7, with: (A.) ED 1. The macrophages adjacent to the implant stain positive for ED 1 (arrow heads). No ED 1 positive cells are observed in the polymer film. (B.) ED 2. The macrophages adjacent to the implant do not stain positive for ED 2. The cells in the lower left corner stain false positive for ED 2. This is due to endogenous peroxidase activity of neutrophilic granulocytes. (C.) ED 3. The macrophages adjacent to the implant do not stain positive for ED 3. Bar indicates 40 µm.

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Fig. 47.5. “Combi” PLLA film: P 7 days after implantation. Note the difference in the intensity of the inflammatory response between the porous and the nonporous side. Bar indicates 40 µm.

subcutaneous tissue under the scar of the sham operation. There are almost no cells localized in the pores of ePTFE (Fig. 47.7A). In contrast, the inflammatory response against porous PLLA was localized mainly in the large pores (Fig. 47.7B). Relative Polymer Surface Area in Sections The relative polymer surface area (RPSA) in sections was determined after in vivo and in vitro procedures. After harvesting, samples were immediately fixed by immersion in a 0.1 M Na-cacodylate buffer, pH 7.4, containing 2% glutaraldehyde and 0.1 M sucrose, for at least 24 h at 4°C. After rinsing with buffer, the samples were dehydrated in graded ethanol series. The samples were then embedded in a position which allowed sections to be cut perpendicular to the longitudinal axis of the polymer films. After embedding in glycolmethacrylate (GMA) (Technovit, Kulzer, Germany), sections of 3 µm were cut on a Jung 1140 autocut microtome, using a D knife with a tungsten carbide cutting edge. The sections were mounted on glass and stained with toluidine blue and alkaline fuchsin.

The relative polymer surface area in sections (RPSA) was determined by morphometrical analysis using light microscopical sections and a Quantimet 520 (Cambridge Instruments) image analyzer. The RPSA was defined as the ratio between the polymer surface area and measurement frame area. The boundaries of the measurement frame were set at the outer boundaries of the polymer surface area. The morphometrical analysis was performed on porous and “combi” films only, because the nonporous film was not eroded nor did it become porous, not even after an immersion or implantation period of 180 days. The RPSA of the nonporous film remained 100%. For each implantation period or immersion period ten measurements of the remaining polymer surface area (of at least two sections) were performed. The mean value and standard deviation were calculated from these ten values. All data were normalized to the initial value (t = 0). The results of the morphometrical analyses are presented in Figures 47.8A (in vitro) and 47.8B (in vivo). The RPSA in vivo shows no significant decrease for the nonporous, the porous or the “combi” film. Also, no decrease was

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Fig. 47.6. Nonporous PLLA film: P 14 days after subcutaneous implantation. Note the large concentration of giant cells (arrow heads). Bar denotes 40 µm.

observed in vitro for the nonporous and porous film up till day 180 and for the “combi” film, up till day 90. It was not possible to obtain the surface area of the “combi” film at day 180, due to fragmentation. During the in vivo experiment with the nonporous polymer film, no erosion or pore formation was observed. Moreover, after immersion in PBS (in vitro), it was not possible to carry the nonporous film through the embedding procedures for light microscopy from day 40 on, due to increased brittleness. The surface area of the remaining nonporous film remained 100% over the entire test period. Mw and Mn Mw (molecular weight) and MN in vitro were determined by gel permeation chromatography (GPC) at 20°C on a Waters Associates GPC apparatus using Waters Associates columns (bead size of 105, 104, 103 Å). A precolumn with a pore size of 500 Å was used. A Waters Associates R 403 differential refractometer was used as a detector. A

sample of 5-10 mg was dissolved in 10 ml chloroform and filtered (Spartan 13/20 filter, 0.45 µl). The injection volume per measurement was 200 µl. Chloroform was used as eluent at a flow rate of 2.0 ml/min. The Mw, Mn and polydispersity ratio were calculated using the calibration data of polystyrene standards of narrow molecular weight distribution dissolved in tetrahydrofuran (THF). The Mw (Fig. 47.9A) and Mn (Fig. 47.9B) decreased for all three films at approximately the same rate. From day 7 on, the porous film generally retained the highest Mw and Mn. The initial (t = 0) Mw of the “combi” film was the highest (167,000) and of the nonporous film the lowest (98,000; porous was 109,000). Before day 7, the curves of the porous and “combi” film showed large fluctuations for Mw. Thereafter, all values for Mw demonstrated a decreasing trend till the 170th day (Mw nonporous: 24,000, porous: 60,000, “combi”:29,000). The Mn curves roughly approximate the same trend as the Mw curves. The fluctuations of the values for the “combi” film are larger.

The Influence of Porosity and Surface Roughness on Biocompatibility

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Fig. 47.7. (A, left.) Porous e-PTFE (P) and (B, right.) porous PLLA (P) after 180 days of implantation. There are almost no inflammatory cells localized in the pores of the e-PTFE film. In contrast, the inflammatory response is localized mainly in the pores F=fibrous encapsulation. Bar indicates 63µm.

Interpretation of Results Inflammatory Response and Biocompatibility The results demonstrate that there are two phases in the inflammatory response against the films. Phase 1 is observed upon implantation of the film. It is mainly caused by the injury sustained by the implantation procedure. This uncomplicated inflammatory response, part of the wound healing reaction, ends after 7-10 days and has been well described.2,20 In this phase, the contribution of the implanted polymer film to the intensity of the inflammatory response is in most cases minimal, except when leakage of large quantities of toxic products occurs.5 After one week, any remaining inflammatory response can be considered as a tissue reaction against the implanted biomaterial.11 This chronic inflammatory response (phase 2) is often described as a foreign body reaction.1,2,5 It mainly consists of macrophages and giant cells surrounding the implant. A minimal inflammatory response is preferred when biomaterials are implanted for a long time span, because a persistent (chronic) inflammatory response may predispose for amyloidosis,21,22 or carcinogenesis.23 A persistent inflammatory response increases the concentration of both serum amyloid A (SAA) and amyloid enhancing factor (AEF) in

blood. SAA may then be converted into amyloid A, which is deposited in tissues. However, the exact mechanism is yet to be fully uncovered, and the relation between the chronic inflammatory response against biomaterials and amyloidosis also remains to be investigated. The Role of Macrophages and Giant Cells ED 1 stains an intracellular antigen, probably associated with the lysosomal membrane. The antigen probably remains present even in the event of fusion of macrophages and formation of foreign body giant cells, since ED 1 positive cells are observed on the same location of foreign body giant cells in conventional light microscopical sections. ED 2 is a marker which is found on mature tissue macrophages. It takes about one week for monocytes/macrophages to express ED 2 on their cell surface after leaving the vascular system. The role of ED 2 and ED 3 in the inflammatory response is not well understood. The different staining patterns of the ED antibodies indicate the possibility of either different populations of macrophages which play a role in the tissue reaction against biomaterials, or macrophages expressing different surface receptors in the course of the inflammatory response against a biomaterial.

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Fig. 47.8. Remaining polymer surface area (RPSA) of the PLLA films in a light microscopical section. (A.) In vitro: RPSA as a function of immersion time. (B.) In vivo: RPSA. Symbols represent: ❑ = nonporous, + = porous, ❍ = “combi”. Bars indicate the standard deviation.

The exact mechanism of foreign body giant cell formations is still to be elucidated. Certain cell types, e.g., T lymphocytes, may play a pivotal role in this process. T helper lymphocytes capable of secreting interferon-γ, especially, may play a role in the fusion of macrophages, forming giant cells.24,25 Activation of macrophages may be induced by various pathways. In one pathway, (limited) damage to neutrophils and macrophages may lead to secretion of cytokines (IL-1), activating T helper lymphocytes. Another possibility is change of shape when a macrophage comes in contact with a boimaterial,26,27 especially when a single macrophage is not able to phagocytose the polymer(fragment). Porosity, Wettability, Degradability and Inflammatory Response Porous PLLA provokes a more intense inflammatory response from day 7 on, despite a higher degradation rate of

the nonporous PLLA in an aqueous environment. This indicates that porosity is an important factor determining the intensity of the inflammatory response against implanted PLLA films. The small difference in the intensity of the inflammatory response between PTFE and ePTFE as compared to porous and nonporous PLLA respectively indicates that porosity as a single factor is not enough to enhance the inflammatory response. The relatively high wettability of PLLA compared to ePTFE allows for better ingrowth of tissue into the pores of PLLA and probably more exposure to other factors, determining the intensity of the inflammatory response. According to many authors, the surface properties of an implant have a large influence on the inflammatory response. Thus, the difference in wettability as one of the surface properties may also be a factor, although Baier et al could not demonstrate a difference in the inflammatory response against smooth metal pieces having a different wettability.28

The Influence of Porosity and Surface Roughness on Biocompatibility

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Fig. 47.9. Molecular weight of the PLLA films as a function of immersion time. (A.) MW. (B.) Mn. Symbols represent: ❑ = nonporous, + = porous, ❍ = “combi”. Standard deviation did not exceed 15%.

However, other authors demonstrated that different polymers induce a different level of IL-1 production, correlating with the intensity of the inflammatory response,29,30,31 although the relation to wettability was not investigated. However, the production of IL-1 seems to be less when using relatively hydrophobic materials such as silicon. Further investigation is needed to establish the precise relation between wettability and inflammatory response. The differences in inflammatory response against the PLLA (degradable) and PTFE (nondegradable) films became apparent from day 40. One reason may be the difference in the rate of degradation and subsequently the difference in the release of degradation products such as monomers, oligomers and finally fragments. As stated previously, only a few (mostly nonporous) PLLA films were observed to be broken into two or three pieces. However, there is probably also increase in surface area which could hardly be observed.

Also, the size of the pores of the porous PLLA films has increased, possibly with the same effect as fragmentation. In the case of PLLA films, the increase in surface area during the degradation process as a single factor may be sufficient for increasing the intensity of the inflammatory response. It can be concluded that biodegradable PLLA films provoke a more intense inflammatory response than nondegradable PTFE films. Also, porosity enhances the inflammatory response. However, porosity enhances the inflammatory response only when the wettability of a biomaterial permits cellular ingrowth. Relative Polymer Surface Area The RPSA, which was only determined for the porous and the nonporous polymer film, did not decrease significantly, neither in vitro nor in vivo. However, these results must be interpreted with some caution, because RPSA

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measurements cannot detect release of small amounts of degradation products from the core of the polymer film. The loss of weight of the nonporous film could not be detected with RPSA measurements since there was no detectable loss of surface area in cross sections of the PLLA film. There was also no indication of measurable surface erosion, such as increasing surface roughness and/or decreasing thickness of the nonporous film. However, this method might be useful in quantifying the later stages of the degradation and the resorption process in vivo, when weight measurement has become practically impossible. The results indicate that in vitro the core of the PLLA films remain largely intact till day 180, and that in vivo there was no resorption of large amounts of polymer. These findings support the results of the weight measurements. Molecular Weight The results of the molecular weight measurements indicate a higher degradation rate for the nonporous film compared to the “combi” and porous film, the latter having the lowest degradation rate. The decrease of MW and MN is obvious for all three types of film, although results obtained with GPC must be interpreted with caution when low molecular weight polymers resulting from degradation processes are examined. Therefore, the differences between the films may not be significant, but the higher MW and MN of the porous film compared to the nonporous film, from day 7 till the end of the experiment at day 180, is very suggestive. The initial MW and MN of the different films are not equal (Figs. 47.9A and 47.9B). A possible explanation may be the washing out of a larger part of the low molecular weight fraction into the media, at the extra rinsing step to remove the sodium citrate, of the films with a porous component; therefore higher MW and MN were measured. However, it can be expected that washing out of low molecular weight fractions probably also occurs during immersion of the nonporous film in the buffer. As a consequence, it can be expected that the difference in MW and MN between the films would disappear in the course of the experiment. This was not the case, indicating another process having an effect on MW and MN. At day 180 the MW and MN values of the “combi” film tend to approach the ones of the nonporous film. This trend corresponds with the erosion of the thin nonporous layer of the “combi” film as observed with SEM: The “combi” film was becoming a porous one. Also, the MW and MN show a lot of fluctuation during the first two weeks of the experiment. These fluctuations might be caused by two phenomena having opposite effects on MW and MN: washing out of low molecular weight fractions and molecular chain fragmentation. However, as stated earlier, this remains hypothetical, as the amount of polymer compared to the amount of buffer in which they were placed did not allow for an analysis using GPC. Degradation by Hydrolysis Hydrolysis of the ester bonds is the major mechanism of PLLA degradation.32 Mechanical stress may enhance the degradation process.33 Radiation (UV, IR) or heat is not ex-

Tissue Engineering of Prosthetic Vascular Grafts

pected to contribute very much to the degradation of polyesters in vivo. The role of enzymes in the degradation process in vivo is not clear. Most enzymes in eukaryotic cells are substrate specific. Therefore, a specific three dimensional structure of the substrate (polymer) is required to reach the active center of the enzyme. This is not likely at 37°C under physiologic enzyme concentrations. The probability is even smaller for the crystalline parts of the polymer (film). However, enhancement of the degradation rate in vitro by enzymes was observed for some polymer/enzyme systems. The role of enzymes may be larger when smaller molecules have been formed by other degradation mechanisms (hydrolysis). This was not investigated in the experiments described in this paper. Degradation of polyesters, such as PLLA, primarily takes place in the amorphous part of the polymer film.34,35 This may explain the increasing crystallinity, also observed by other authors.34 However, in order to explain the difference in degradation rate between the nonporous and porous PLLA film, one must assume other processes taking place during degradation. Vert et al, using different polyester “plates” of 2 x 17 x 20 mm, observed the formation of a polymer layer around the core of the polymer film, largely preventing degradation products (e.g., oligomers) released in the core, to diffuse freely to the aqueous environment. The chemical reactive endgroups of the accumulated oligomers in the core then enhance the degradation (autocatalytic) process. The nonporous polymer film is more susceptible to this form of degradation, probably because a larger inner compartment can be formed compared to the porous and “combi” film. The surface/volume ratio of the latter two are probably larger than the cut off value of the polymer surface/volume ratio regarding this effect. However, other mechanisms are also possible. There is less flow of the aqueous media in pores of the porous polymer film. Therefore, the concentration of oligomers might be higher in the pores, enhancing degradation. Nevertheless, this mechanism apparently has less effect compared to the mechanism(s) leading to a higher degradation rate of the nonporous PLLA film.

Summary and Conclusions In summary, it can be stated that the parameters of porosity and surface roughness do influence the ultimate biocompatibility of implanted biomaterials. Potential degradation of a biomaterial will enhance these reactions. Porosity A high porosity will, as a rule, induce a higher inflammatory response as compared to smooth and solid biomaterials. The size of pores will, in turn, determine the type of foreign body reaction. Small (surface) pores (up to 10 microns) will give rise to more adhesion. In addition, guidance of cells may occur. Ingrowth will not take place. Larger pores will facilitate tissue ingrowth and therewith provide a mechanical anchorage of tissue to biomaterial. Optimal pore sizes have been estimated to be between 50 and 200 microns.

The Influence of Porosity and Surface Roughness on Biocompatibility

Surface Roughness As a rule, a high surface roughness will induce a more severe foreign body response and also provide better anchorage of cells to a biomaterial. The form of the roughness will determine if other aspects are also involved, e.g., grooved substrata36 will, depending on their dimensions and orientation, improve cell proliferation and orientation. Degradation and Biocompatibility In general, a degrading biomaterial will evoke a more severe inflammatory response as compared to a nondegradable material. If this degradable material is porous, we have demonstrated that, in the case of, e.g., collagen, the rate of degradation (and thus the degree of foreign body response) is higher, whereas in the case of polyesters as lactic acid, a porous form will degrade more slowly than a solid form. In conclusion we can state that porosity and roughness will influence cellular adhesion and the rate of foreign body reaction. However, parameters such as biodegradability will have their own effects on the ultimate degree of biocompatibility. References 1. Ratner BD. Biomedical applications of synthetic polymers. In: Aggarawal SL, ed. Comprehensive Polymer Science. Oxford, U.K.: Pergamon Press, 1989:201-247. 2. Spector M, Cease C, Tong-Li X. The local tissue response to biomaterials. Crit Rev Biocompatability 1989; 5:269-295. 3. Black J. Systemic effects of biomaterials. Biomaterials 1984; 5:11-18. 4. Marchant RE, Anderson JM, Dillingham EO. In vivo biocompatability studies VII, Inflammatory response to polythylene and to cytotoxic polyvinylchloride. J Biomed Mater Res 1986; 20:37-50. 5. Wachem PB, Luyn MJA, Nieuwenhuis P, Koerten HK, Olde Damink L, Hoopen H ten, Feijen J. In vivo degradation of processed dermal sheep collagen evaluated with transmission electron microscopy. Biomaterials 1991; 12:215-223. 6. Schakenraad JM, Busscher HJ, Wildevuur CHR, Arends J. The influence of substratum free energy on growth and spreading of human fibroblasts in the presence and absence of serum proteins. J Biomed Mater Res 1986; 20:773-784. 7. Matlaga BF, Yasenchak LP, Salthouse TN. Tissue response to implanted polymers: The significance of shape. J Biomed Mater Res 1976; 10:391-397. 8. White RA, Hirose FM, Sproat RW, Lawrence RS, Nelson RJ. Histopathologic observations after short term implantation of two porous elastomers in dogs. Biomaterials 1981; 2:171-176. 9. Cheroudi B, Gould TRL, Brunette DM. Titanium covered micromachined grooves of different dimensions affect epithelial and connective tissue cells differently in vivo. J Biomed Mater Res 1990; 24:1203-1219. 10. Bakker D, Blitterswijk CA van, Hesseling SC, Grote JJ, Daems WT. Effect of implantation site on phagocyte/polymer interaction and fibrous capsule formation. Biomaterials 1988; 9:14-23. 11. Schakenraad JM, Nieuwenhuis P, Molenaar I, Helder J, Dijkstra PJ, Feijen J. In vivo and in vitro degradation of

529 glycine/DL-lactic acid copolymers. J Biomed Mater Res 1989; 23:1271-1288. 12. Rozema FR, Bruin WC de, Bos RRM, Boering G, Nijenhuis AJ, Pennings AJ. Late tissue response to bone plates and screws of poly L-lactide used for fracture fixation of the zygomatic bone. In: Doherty PJ, Williams RL, Williams DF, eds. Advances in Biomaterials, BiomaterialTissue Interfaces. Amsterdam: Elsevier, 1992:349-355. 13. Schakenraad JM, Oosterbaan JA, Nieuwenhuis P, Molenaar I, Olijslager J, Potman W, Eenink MJD, Feijin J. Biodegradable hollow fibers for the controlled release of drugs. Biomaterials 1988; 9:116-120. 14. Spector M. Historical review of porous coated implants. J Arthroplasty 1987; 2:163-177. 15. Baier RE. Applied chemistry at protein interfaces. Adv Chem Ser 1975; 145:1-25. 16. Wachem PB, Beugling T, Feijen J, Bantjes A, Detmers JP, Aken WG van. Interactions of cultured human endothelial cells with polymeric surfaces of different wettabilities. Biomaterials 1985; 6:403-408. 17. Schindler A, Harper D. Polylactide 2, viscosity molecular weight relationship and unperturbed chain dimensions. J Polymer Sci 1979; 17:2593-2599. 18. Busscher HJ, Pelt AWJ van, Jong HP de, Arends J. Effect of spreading pressure on surface free energy determinations by means of contact angle measurements. J Colloid Interfacial Sci 1983; 95:23-27. 19. Blaauw EH, Jonkman MF, Gerrits PO. A rapid connective tissue stain for glycol methacrylate embedded tissue. Act Morphol Neerl-Scand 1987; 25:167-172. 20. Cotran RS, Kumar V, Robbins SL. Inflammation and repair. In: Cotra RS, Jumar V, Robbins SL, eds. Pathologigal basis of diseases. Philadelphia: Saunders, 1994:51-92. 21. Picken MM, Gallo GR. Ameloid enhacing factor and inflammatory reaction. Lab Invest 1990; 63:586-587. 22. Kisilevski R. Amyloidosis, In: Rubin E, Farber JL, Lippincott JB, eds. Pathology. Philadelphia:, 1994:1163-1174. 23. Weitzman SA, Gordon LI. Inflammation and Cancer: Role of phagocyte-generated oxygen carcinogenesis. Blood 1990: 76:655-663. 24. Mentzer SJ, Valler DV, Burakoff SJ. Gamma-interferon induction of LFA-1 mediated homotypic adhesion of human monocytes. J Immunol 1986; 137:108-113. 25. Most J, Neumayer HP, Dierich MP. Cytokine-induced generation of multinucleated giant cells in vitro requires gamma-interferon and expression of FLA-1. Eur J Immunol 1990; 20:1661-1667. 26. Shaw LM, Messier JM, Mercurio AM. The activation dependent adhesion of macrophages to laminin involves cytoskeletal anchoring and phosphorylation of the alpha6 beta-1 integrin. J Cell Biology 1990; 110:2167-2174. 27. Ingber DE, Folkman J. Tension and compression as basic determinants of cell form and function: Utilization of a cellular tensigrity mechanism. In: Cell Shape: determinants, regulation and regulatory role. New York: Academic Press, 1989:3-31. 28. Baier RE, Meyer AE, Natiella JR, Natiella RR, Carter JM. Surface properties determine bioadhesive outcome: Methods and results. J Biomed Mater Res 1984; 18:337-355. 29. Miller KM, Anderson JM. Human monocyte/macrophage activation and interleukin 1 generation of biomedical polymers. J Biomed Mater Res 1988; 22:713-731. 30. Miller KM, Rose-Caprara V, Anderson JM. Generation of IL-1 like activity in response to biomedical polymer

530 implants: A comparison of in vitro and in vivo models. J Biomed Mater Res 1989; 23:1007-1026. 31. Krause TJ, Robertson FM, Liesch JB, Wasserman AJ, Greco RS. Differential production of IL-1 on the surface of biomaterials. Arch Sorg 1990; 125:1158-1160. 32. Richardson MJ. Thermal analysis. In: Booth C, Price C, eds. Comprehensive Polymer Science. Oxford: Pergamon Press, 1989:867-901. 33. Jamshidi K, Hyon S-H, Nakamura T, Ikada Y, Teramatsu Y. In vitro and in vivo degradation of poly-L-lactide fibers. In: Christel P, Neunier A, Lee AJC, eds. Advances in biomaterials: Biological and biomechanical perfor-

Tissue Engineering of Prosthetic Vascular Grafts mances of biomaterials. Amsterdam: Elsevier, 1986:227-233. 34. Leenslag JW, Pennings AJ, Bos RRM, Rozema FR, Boering G. Resorbable materials of poly-L-lactide VII: In vivo and in vitro degradation. Biomaterials 1987; 8:311-314. 35. Schakenraad JM, Hardonk MJ, Feijen J, Molenaar I, Nieuwenhuis P. Enzymatic activity towards poly (L-lactic acid) implants. J Biomed Mater Res 1990; 24:529-545. 36. Braber ET den, Ruitjer JE de, Smits HTJ, Ginsel LA, Recum AF von, Jansen JA. Quantitative analysis of cell proliferation and orientation on substrata with uniform parallel surfaxe micro-grooves. Biomaterials 1996; 17:1093-1099.

Scaffold Engineering Surface Modification

CHAPTER 48 Microgroove Driven Tissue Ingrowth Edwin T. Den Braber, John A. Jansen

Implants, Tissue Engineering, and Biomaterials

D

uring the last two decades the availability and application of medical implants has in creased dramatically. This concerns a broad variety of medical implants ranging from knee prostheses to insulin infusion pumps, and from vascular grafts to pacemakers. Some estimate figures were presented by Ratner1 in his Presidential Address for the Society for Biomaterials in 1993 (Table 48.1). Although he emphasized that these figures were estimates, it is clear that the use of implants is considerable. Long term projections even suggest that implant applications are going to rise in the future. Factors that contribute to this increase can be ascribed roughly to three major causes. First, the life expectancy of humans increases. This will inevitably lead to a rise in the demand for implants like hip replacements or artificial lenses for the treatment of geriatric diseases and defects. Second, more and more medical treatments are going to include the use of implants in the future. One example of such a development is the use of percutaneous implants in dialysis.2 Instead of treating patients with chronic renal failure though intermittent hemodialysis, percutaneous implants enable continuous ambulatory peritoneal dialysis (CAPD). Third, technologies are evolving that open new ways of treating specific disorders or defects. This is demonstrated by techniques that are being developed in the field of tissue engineering.3 Basically, tissue engineering combines the principles and methods of the life sciences with those of engineering to elucidate fundamental understanding of structure-function relationships in normal and diseased tissues, to develop materials and methods to repair damaged or diseased tissues, and to create entire tissue replacements.4 Tissue engineering thus spans controlling cellular responses to implant materials, manipulating the healing environment to control the structure of the regenerated tissue, producing cells and tissues for transplantation into the body, and developing a quantitative understanding of many biological equilibrium and rate processes.5 Biomaterials play an important role in many of these activities. Originally, inertness was thought to be one of the major contributions of the performance of an implant or biomaterial, but later Williams adapted the definition of biocompatibility to include the idea that a biomaterial perform with an appropriate host response in a specific application.6 Sadly, most currently used implant materials do not possess these desirable qualities. At his Presidential Address, Ratner1 voiced this problem very vividly by saying, “For the majority of our widely used bio-materials, no one sat down in advance and said ‘How can I engineer the surface of this material to produce the desired biological response?’” He stressed that most currently used biomaterials, although demonstrating generally satisfactory clinical performance, were “ad-hoc” biomaterials, developed through a trial and error optimization, rather than being engineered to produce, for instance, a desired interfacial interaction. The Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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Table 48.1. Selected biomedical implant applications; magnitude of usea Application

Numbers Used per Year

Ophthalmologic Intraocular lenses Contact lenses Retinal surgery implants Prothesis after enucleation

1,400,000 250,000,000b 50,000 5,000

Cardiovascular Vascular grafts Arteriovenous shunts Heart valves Pacemakers Blood bags Reconstructive Breast protheses Nose, chin Penile Dental

350,000 150,000 75,000 130,000 30,000,000 100,000 10,000 40,000 200,000

Orthopedic Hips Knees Shoulders, finger joints

90,000 65,000 50,000

Other devices Ventricular shunts Catheters Oxygenerators Renal dialysers Wound drains Sutures

21,500 200,000,000 500,000 16,000,000 3,000,000 20,000,000

a b

Approximate annual usage in United States of America Worldwide

importance of the issue raised by Ratner is underlined by recent estimates that indicate that biomaterials such as metals, ceramics, and polymers are found in more than 5,000 different medical devices and almost 40,000 different pharmaceutical products, with a collective annual sales approaching 100 billion US dollars.5 If we include all products that utilize biomaterials, then the number of people affected worldwide exceeds 1000 million per year.4,7 As a possible handle for the development of “smart” biomaterials, Ratner1 advocated an engineering approach to achieve the integration between the biomaterial and the surrounding tissues. He suggested that it is worthwhile to look at how nature handles specificity and rapid-reaction kinetics, keeping in mind that three design themes are iterated and exploited throughout biology, i.e., order, recognition, and mobility. However, as Brunette8 later remarked, using the principles of design to achieve the optimal structure differs from attempting to reproduce the original structure. According to Brunette, attempting to imitate nature is approximate since the complexity and variability of natural systems make them impossible to replicate with perfect fidelity. This has led to the production of structures and

devices that closely resemble the original structure, but are in no way the best replacement for the original biological structure. The development of these kinds of structures caused Ratner to state during his Presidential Address1 that “existing biomaterials, although demonstrating generally acceptable clinical success, look like dinosaurs poised for extinction in the light of the winds of change blowing through the biomedical, biotechnological, and physical sciences”.

Surface Micropatterns Manipulating Cellular Behavior Among the tools that Ratner identified as having a great potential for creating engineered biomaterials is the use of nano- and micropatterned surfaces.1 The possibility to influence the behavior of cells by the topographical morphology of the surface on which these cells are cultured was first discovered during the first part of this century. In 1914, Ross G. Harrison of the Osborn Zoological Laboratory of Yale University cultured cells on spiderwebs.9 In retrospect, this in itself seems quite an achievement, especially when we keep in mind that in 1914 investigators did not have all the technical equipment that are available in modern cell culture laboratories nowadays. During his studies, Harrison noticed that the direction of movement of the cells was influenced by the structure of the fragile substrata on which these cells were incubated. Furthermore, he noticed that the cells adapted a shape that seemed to be governed by the linear spiderweb threads. Later, this phenomenon was confirmed by Loeb and Fleisher,10 who introduced the term stereotropism, which was described as the direction in which cells move, governed mainly by the contact with solids or very viscid bodies like fibers or fibrin. In 1945, Weiss conducted experiments with a wide range of substrata like plasma clots, fish scales, and engraved glass, and termed the changes in cell behavior observed by Harrison “contact guidance”.11 Surprisingly, no further attention was paid to this guidance phenomenon until the early 1970s. It was Rovensky et al12-13 and Maroudas14-15 who rediscovered that the behavior of cells is affected by the topography of a substratum surface. Rovensky12-13 studied the behavior of fibroblast-like cells on different kinds of substratum surfaces with an orderly distribution of 40.0 µm deep grooves having a triangular profile. He reported that the cells had a bipolar elongated shape, were orientated after adhesion, and grew parallel to the grooves. Meanwhile, Maroudas14-15 studied the growth of fibroblasts on small glass beads (diameter 20-60 µm), fibers, and platelets. He observed that cells grown on beads with a large diameter tended toward forming multilayers, while smaller beads progressively failed to support growth. Since these studies by Rovensky and Maroudas, several investigators have studied the behavior of various types of cells to a variety of microtextured substrata materials. These studies have been reviewed very recently by Singhvi et al,16 von Recum and van Kooten,17 and Brunette.8 In these in vitro and in vivo studies, some specific alterations in cellular behavior were seen, while other changes that were suggested by several leading researchers in this field were not. In the following paragraphs we will attempt to give a com-

Microgroove Driven Tissue Ingrowth

prehensive listing of the most obvious and general accepted alterations in cell behavior as a result of cells contacting surfaces with a pattern of parallel microgrooves. We will demonstrate this by describing various studies that were performed at our laboratories. In addition, we will discuss the leading theories that attempt to create a model in which the phenomenon of contact guidance is clarified.

Changed Cell Shape and Cellular Orientation If cells are cultured on a surface with parallel microgrooves as shown in Figure 48.1, the first thing that can be observed is the change of the shape of the cells. In earlier studies,18-19 we investigated several groove and ridge dimensions to study and quantify the influence of these groove patterns on the size, shape, and orientation of the cells cultured on these substrata. In order to obtain a microgrooved substratum, we first produced silicon oxide molds in a class 100 clean room using photolithography.20-21 Photolithography is a technique that is developed for the production of microelectronic components like, for example, computer microcircuits. In our experiments, these mold surfaces were produced by coating silicon oxide masks with high reflective chrome, after which the chrome was coated with a thin (0.5 µm) layer of positive photoresist. Subsequently, the photoresist was exposed and developed, uncovering the underlying chrome, which was etched. Finally, the unexposed photoresist was stripped off, thus creating a parallel groove pattern. The substrata were finally obtained by covering the molds with polydimethylsiloxane, thus producing negative surface replicas which possessed either a smooth surface or a surface with parallel grooves, showing a wide variety of groove and ridge widths (Table 48.2). During our experiments, the groove depth was either 0.5 µm or 1.0 µm. After polymerization, the silicone rubber castings were removed, cut into their appropriate

533

circular shape and size, washed, sterilized, and inspected as described elsewhere.18-19 For the evaluation of the cellular behavior on the smooth and microtextured silicone substrata, primary culture rat dermal fibroblasts (RDFs) were used. These cells were harvested from ventral skin grafts taken from male Wistar rats (100-120 g). After dissociation, these cells were incubated for several days according to a specific protocol.18-19 Subsequently, the fifth generation was used for incubation on the microtextured surfaces. Therefore, the smooth and microtextured substrata were placed in the culture wells of 24 well plates, after which approximately 1.0 x 104 viable RDFs ml-1 were added to each substratum. The RDFs were incubated on a specific substratum up to 7 days under static conditions. The effect of the surface microgeometry on the cellular morphology was quantified by digital image analysis.18-19

Table 48.2.Dimensions of the micro features on the substrata surfaces (Gd=groove depth, Gw=groove width, Rw=ridge width, and P=pitch). Gd (µ µm)

Gw (µ µm)

Rw (µ µm)

P (µ µm)

1.00 1.00 1.00 1.00 1.00 1.00 0.45 0.45 0.45

1.00 1.00 1.00 1.00 4.00 8.00 2.00 5.00 10.00

1.00 2.00 4.00 8.00 1.00 1.00 2.00 5.00 10.00

2.00 3.00 5.00 9.00 5.00 9.00 4.00 10.00 20.00

Fig. 48.1. Three dimensional representation of the results of AFM measurements on a microtextured substratum surface with parallel grooves of 5.0 µ m, separated by 5.0 µm ridges. The grooves are approximately 0.5 µm deep. Different X- and Y-axis magnifications were used to clarify the conformation of the substratum surface.

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Fig. 48.2. Schematic representation of a rat dermal fibroblast (RDF) on a microtextured substratum. The parameters measured during digital image analysis (DIA) were the RDF surface area (area within perimeter), the longest length of the cell (L), the cellular breadth (B), perimeter (P), circularity (not shown), angle of cellular orientation (α ), and the number of pitches spanned by the cell (N).

In short, RDFs were photographed by phase contrast microscopy on every day of the incubation period. These photographs were scanned digitally and analyzed with an image analysis program. This program made it possible to measure several cell parameters, i.e., the cellular surface area, cellular perimeter, cellular circularity, maximum cell length, cell breadth perpendicular to the maximum length, the angle of cellular orientation relative to the surface grooves (α), and number of pitches spanned by a single cell (Fig. 48.2). These studies provided a lot of interesting data, but most of all clearly showed that the RDFs on surfaces with a ridge width smaller than 4.0 µm were highly orientated (α < 10°) and elongated along the surface grooves (Figs. 48.3 and 48.4). However, if the ridge had a width larger than 4.0 µm, then the cellular orientation was random (≈45°) and the shape of the RDFs became significantly more circular (Fig. 48.5). The appearance and orientation of the cells on the surfaces with ridge widths larger than 4.0 µm in many cases proved to be not significantly different from those on the smooth substratum surfaces (Fig. 48.6). Furthermore, it became apparent that the ridge width was the most important parameter, since varying the groove width and groove depth did not affect the RDF size, shape, or the angle of cellular orientation (α) significantly.

Cell Attachment on Microtextured Surfaces During these studies, we found after careful examination of the phase contrast images that the fibroblasts showed indications of attaching specifically to the ridges of the surface patterns (Fig. 48.4). If the fibroblasts did actually attach solely to the surface ridges, this would mean that the cells displayed a specific preference of attachment location. Intrigued by the phase contrast microscopy/digital image analysis results, we performed experiments which were designed to show us more of the intricate interaction between

the microtextured substratum surfaces and the overlaying fibroblasts. During these additional experiments, we cultured the RDFs similarly as in the earlier studies,18-19 but visualized specific cytoskeletal components and proteins interacting in the attachment of the RDFs with the following monoclonal and polyclonal antibodies:22 1. The mouse monoclonal antibody hVIN-1, specific for vinculin;23 2. The rabbit monoclonal anti-fibronectin antibody FN-3E2;24 3. A rabbit polyclonal antiserum raised against bovine fibronectin;25 4. The rabbit polyclonal antiserum against bovine vitronectin;26 5. A polyclonal antiserum raised against human vitronectin in rabbits, which has been shown to have a good crossreactivity with rat vitronectin, but not with bovine vitronectin.27 After incubation, these primary antibodies were incubated with the appropriate secondary antibody, i.e., fluorescein isothiocyanate (FITC)-conjugated goat anti-mouse IgG or FITC-conjugated goat anti-rabbit IgG. For the staining of the RDF filamentous actin no antibodies were used, since F-actin was visualized with thiorhodamine isothiocyanate (TRITC)-labeled phalloidin. Immediately after performing the double stain procedures, the RDFs were examined with a confocal laser scanning microscope (CLSM), which was equipped with a krypton/argon mixed gas laser. This type of laser offers separate, well spaced wavelengths for the excitation of FITC (λ = 488 nm) and TRITC (λ = 568 nm).28 Next to the fluorescence mode, the reflection mode of the CLSM was used to visualize the underlying substratum surface.22 The resulting digital images were captured and overlay images were created, thus making it possible to capture the fluorescent and reflection data in one 24 bit RGB

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Fig. 48.3. Phase contrast image of RDFs on a B substratum (Gw 1.0 µm, Rw 1.0 µm, Gd 1.0 µm, bar = 100 µm) on day 1. The cells are highly aligned and elongated along the surface grooves.

Fig. 48.4. Phase contrast image of RDFs on a substratum with grooves of 8.0 µm wide and 1.0 µm deep, while the ridge has a width of 1.0 µm (bar = 100 µm) after 2 days of incubation. These substrata are a negative replica of the substrata in Figure 48.5. The RDFs are clearly orientated. Cell protrusions attach to the ridges (arrows ).

(Red-Green-Blue) picture. Creation of these 24 bit images enabled the composition of 1 digital image with 3 different information levels, each within one color segment, and offered the possibility to investigate the stained objects in conjunction with each other and the surface patterns. First, the (acute) angle of orientation of F-actin, vinculin, fibronectin, vitronectin, and the surface grooves relative to a virtual X-Y axis were measured. Second, the relative position and the angle of the linear components of vinculin, fibronectin, and vitronectin were compared with the position and angle of orientation of the actin filaments. Finally, the location of

vinculin relative to the grooves and ridges of the microtextured surface was charted and analyzed.22 The results of the CLSM observations and additional image and statistical analysis showed that the microfilaments and vinculin aggregates of the RDFs on the 2.0 µm grooved substrata were orientated along the surface grooves, while these proteins were significantly less orientated on the 5.0 and 10.0 µm grooved surfaces (Fig. 48.7). In contrast, bovine and endogenous fibronectin and vitronectin were orientated along the surface grooves on all textured surfaces. These extracellular proteins did not seem to be hindered by

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Fig. 48.5. Phase contrast image of RDFs on a substratum with grooves of 1.0 µm wide and 1.0 µm deep, and 8.0 µm wide ridges (bar = 100 µm) on day 2. Appearance of and orientation of the fibroblasts does not differ significantly from the RDFs on a smooth substratum surface.

Fig. 48.6. Phase contrast image of RDFs on a smooth substratum after 2 days of incubation (bar = 100µm). The well spread RDFs have a characteristic multipolar appearance and are randomly orientated.

the surface grooves, since many groove spanning filamentous deposits were found on all microgrooved surfaces (Fig. 48.8). Vinculin was located mainly on the surface ridges on all textured surfaces (Fig. 48.9). These findings were later confirmed in additional studies.29-30 During these investigations we made ultrathin sections of RDFs cultured on silicone and titanium microgrooved surfaces with parallel grooves ranging from 1.0 µm up to 10.0 µm. The depth of these grooves varied between 0.45 µm and 2.2 µm. Transmission electron microscopy (TEM) again showed that the focal adhesion points of the RDFs on both the silicone rubber and titanium microtextured surfaces were located mainly on the surface ridges (Fig. 48.10). On the titanium surfaces, TEM also re-

vealed that these focal adhesion points were occasionally wrapped around the edges of the ridges. Attachment of the cells on the silicone rubber and titanium microtextured substrata was never observed on the surfaces with 1.0 or 2.0 µm grooves. Only the RDFs on the 5.0 µm and 10.0 µm grooved surfaces protruded into the grooves, while attachment to the groove floor was observed only on the 10.0 µm textures. In addition, only the RDFs on the titanium 5.0 µm and 10.0 µm grooved surfaces possessed focal adhesion points on the walls of the grooves (Fig. 48.10). Comparison between the observations of the cells on the microtextured silicone rubber and titanium substrata suggested that material specific properties did not influence the orientational effect of the surface texture on the observed RDF cellular behavior.30

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Fig. 48.7. Digital overlay image of a reflection micrograph of a 5.0 µm grooved silicone surface and the corresponding RDF actin after a incubation of 3 days. Fibronectin filaments at the ventral side of this fibroblast can be seen in Figure 48.8. Note the similarity in the orientation of the microfilaments and fibronectin filaments in Figure 48.8.

Fig. 48.8. Immunofluorescence micrograph of a rat fibronectin staining on a surface with 5.0 µm grooves (incubation period of 3 days) showing thin fibronectin filaments at the leading edge of the cell. The main fibronectin filaments possess similar orientational vectors, which are directed roughly parallel to the groove pattern.

Fig. 48.9. Overlay image of the location of vinculin (bright spots) in relation to the 10.0 µ m grooves after an incubation of 5 days. Vinculin is located mainly on the surface ridges and often attaches to the edges of these ridges.

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Fig. 48.10. Transmission electron micrograph of a RDF on a 5.0 µ m grooved titanium surface after 3 days of incubation. The cell protrudes into the groove, but does not contact the bottom of the groove. Focal adhesion points (arrowheads) can be seen on the edge of the ridge and the wall of the surface groove. Larger magnification showed that the focal adhesion point on the lefthand ridge edge was wrapped around this edge.

Other Surface Topography Induced Cell Behavior Alterations In Vitro Apart from changes in cell size, shape, orientation, and attachment, investigators have suggested and studied many other cell processes that could possibly be influenced by the surface topography of substrata. One of these processes is cell proliferation. On the basis of the results of some of our earlier studies,18,31 we concluded that the presence or dimensions of the parallel surface grooves as used in our experiments did not result in a change in RDF proliferation rate. This conclusion was, however, in contrast to the findings of Green et al and Ricci et al. For example, Green et al32 reported that especially abdomen fibroblasts (CCD-969sk) cultured on surfaces with 2.0 and 5.0 µm square pillars showed increased proliferation rates. In addition, Ricci et al33 evaluated the in vitro growth of rat tendon fibroblasts and rat bone marrow colonies on unidirectional (grooved) surface microgeometries. They found that the overall colony growth rate was changed, and concluded that surface microgeometry could be used to control the growth rate at implant surfaces. However, this study by Ricci et al also showed that the response to surface topography is dependent on cell type, which could account for the discrepancy in the results that we found between the proliferation rates of our rat dermal fibroblasts, the results of Ricci et al, and the data of Green et al. A point of discrepancy could be the use of different surface textures. Although Ricci et al used a pattern of parallel grooves, Green et al used a texture consisting of square pillars. This could mean that we not only have to consider the possibility of cell type dependent differences in cell proliferation, but also have to contemplate the effect that different surface textures could trigger. Another process that seems to be affected by the topography of the surface that the cell contacts is that of the cell metabolism. Several publications have reported behavioral differences of cells on microtextured surfaces, such as

changes in cellular differentiation, DNA/RNA transcription, cell metabolism, and cellular protein production.16-17,34-35 For example, increased F-actin content and increased persistence and speed of cell movement of macrophage-like cells on grooved surfaces were observed, and changes in the regulation of fibronectin mRNA levels, mRNA stability, and fibronectin stability and assembly on surfaces with microgrooves have been reported. Similar results were described by Hong and Brunette,36 who saw clear differences in the fibronectin mRNA and proteinase levels between oral epithelial cells that were cultured on smooth and microgrooved surfaces. Consequently, they remarked that these surface topography induced changes could be considered as “good news/bad news”. According to these investigators, the good news was the fact that specific surface topographies could be used to enhance the production of a specific protein, while the bad news might be that the production/secretion of other proteins might also be enhanced. If the production/secretion of these “other proteins” would be elevated, this might have a deleterious effect on the integration of the implant. For example, a rise in the production/release of proteinases could result in a large scale degradation of the connective tissue. This simple example demonstrates that, at the molecular level, the regulation of cell function by altering the surface topography of an implant or substratum could be a very complex affair.

Microtextured Surfaces In Vivo Up to this point, we have only described the effect of microtextured surfaces on cells in vitro. However, the results of these in vitro studies have led to the idea that surface microtexturing could be used deliberately to achieve certain desired end results in processes like morphogenesis, cell invasion, repair, and regeneration.4 If this hypothesis proves to be true, it is obvious that surface texturing can be a very important tool in designing successful implants.17,37

Microgroove Driven Tissue Ingrowth

Sadly, in vivo studies with microtextured implants challenging these ideas are scarce. In addition, review of these in vivo studies shows that the design of the used textured implants is very diverse. But even with this large diversity, it is possible to perceive the possible potential of microtextured implant surfaces on several implant related processes. For example, some studies38-39 have reported on the reduction of epithelial downgrowth with microgrooved skin penetrating devices. Other investigators, who implanted microporous or pillared surfaces subcutaneously, found tightly adherent fibrous capsules without inflammatory cells,40 reduced fibrosis,41 and improved blood vessel proximity.41 In order to investigate the effect of microtextured implants on the surrounding tissue and to be able to compare our in vitro and in vivo data, we implanted the smooth and microtextured silicone rubber substrata with 2.0, 5.0, and 10.0 µm grooves (Table 48.2) subcutaneously in a total of 12 female New Zealand White rabbits for periods of 3, 7, 42, and 84 days.42 Scanning electron microscopy (SEM) showed fibroblasts, erythrocytes, lymphocytes, macrophages, fibrin, and collagen on all implant surfaces after 3 and 7 days (Fig. 48.11). After 42 and 84 days only a little collagen and a small number of fibroblasts, but no inflammatory cells, were seen on the implant surfaces. In contrast with the RDFs in the in vitro experiments, the fibroblasts that were observed on the microtextured implant surfaces, were not orientated along the surface grooves. A possible explanation for these differences between in vitro and in vivo orientational cell behavior could be that the cells that are used in in vitro studies are isolated cells, which have no contact with other cells, cell types, or ECM. Previous studies18,37,43 have shown that prolonged in vitro incubation on microtextured surfaces results in the formation of cell-cell contacts, an increase of the spread area, and a decrease of the orientation of the cells on these surfaces. Consequently, it was supposed that the observed guidance phenomenon is an initial response of cells in vitro to certain microtextured surfaces, which is lost

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gradually after cell-cell contacts are formed.18,37,43 In tissues, these contacts with other cells are already present, which could mean that the orientational effect of the textured surfaces is overruled by stronger tissue related signals or cues. Three dimensional reconstruction of CLSM images and normal light microscopy showed no significant differences between the thickness of the capsule surrounding the smooth and microgrooved implants. Since differences between the 2.0, 5.0, and 10.0 µm grooved implants were not detected, we concluded that the depth of the grooves used was not sufficient to facilitate “mechanical interlocking”. This interlocking would reduce the stress and movement at the implant interface and limit the consequential “mechanical irritation” of the surrounding tissues, which is supposed to induce tissue damage, fibrosis, and severe inflammatory responses.16-17,40-41 Other investigators have indeed reported reduction of the capsule size due to microtextured surfaces. However, review of these studies40-41,44-45 shows that the surface texture of the implants in these studies differs significantly from our implants, both in terms of microfeature appearance (pores, pillars, tapered pits, or V-shaped grooves) and dimensions (feature depth, size, and pitch). As a result, the mechanical irritation of the smooth and grooved surfaces would be comparable, resulting in capsules of equal thickness. Furthermore, our textured implants possessed one smooth and one textured side. Although this opened up the possibility for intraimplant evaluation, it did not enhance possible mechanical interlocking between the implant and the surrounding tissues. Therefore, it can be questioned whether the capsule thickness would have been less if both sides of the implant had been textured. Furthermore, normal light microscopy did show a significantly lower number of inflammatory cells, and a significantly higher number of blood vessels, in the capsules surrounding the microgrooved implants. The cause of these differences remains unclear. However, it is possible to speculate that the differences in the number of blood vessels were

Fig. 48.11. On this SEM image (3 days of implantation), the 2.0 µm grooved silicone surface is visible underneath the collagen fibers. Several punctured erythrocytes (E) and a macrophage (M) were located within the collagen matrix.

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part of the proliferation phase of the wound healing process.46 This proliferation phase is a part of the formation of granulation tissue, which is characterized by high fibroblast densities, the formation of new blood vessels, and a new connective tissue matrix.46 After repair, the number of the vessels decreases generally, marking the end of the wound healing process and the start of a steady state. The fact that more vessels were observed around the textured implants during our study could indicate a higher rate of tissue repair around these implants. Further research is, however, required to determine if this in fact is the case.

The Hypotheses In the previous paragraphs we have presented several studies with microgrooved (implant) surfaces. In spite of the fact that several publications like these and some excellent reviews 3,16-19,31,37 have reported on the effects of microtextured surfaces, little is known about the exact mechanism whereby surface topography exerts its effects. Several theories have been suggested, however. First, it has been hypothesized that wettability plays a role in these phenomena. A microtextured surface could possess local differences in surface free energy which promote a specific deposition pattern of the substratum-bound attachment proteins.16-17,47-49 In addition, the spatial arrangement of the adsorbed proteins50-52 and the conformation of these proteins53-54 would be influenced by these substratum surface properties. Second, it has been suggested that the specific geometrical dimensions of the focal adhesion plaques force a cell on a surface with small grooves and ridges to orientate itself parallel to these ridges.55-56 This hypothesis is based on the observation that a minimum length of 2.0 µm is required for focal contacts to establish adhesion.57 This implies that, if the ridge width increases, multiple vectors of adhesion plaque orientation are possible, enabling less orientated cell attachment. Finally, a third hypothesis58-59 supposes that the orientation and alignment of cells on microtextured surfaces are a part of the cellular efforts to reach a biomechanical equilibrium with the net sum of forces minimized. This phenomenon has, for instance, been described extensively in the so-called tensegrity models.58-60 According to these models, it is possible that the anisotropic geometry of substratum surface grooves and ridges establishes stresses and shear-free planes that influence the direction of microtubule61 and microfilament growth62-63 in order to create a force economic situation. Given the current available information, it is impossible to express an opinion on which or whether one of these hypotheses is true. On the other hand, several separately performed studies have reported that surface microtextrues can have a profound effect on specific elements of the cytoskeleton like, for example, the microfilament bundles,62-65 focal contacts,55 and microtubules,61 making it safe to say that microtextured surfaces influence the orientation of the intracellular and extracellular proteins. Although results of our studies corroborate with all three hypotheses, they do not justify a specific choice for one of these hypotheses. The differences in deposition patterns and the appearances of the ECM proteins during the CLSM study for example, make it possible to suggest that surface properties like surface free

Tissue Engineering of Prosthetic Vascular Grafts

energy have an influence on the displayed cellular behavior. The vinculin location and orientation however, pleas in favor of the “ridge width” theory, although warped focal adhesion points that were found in the TEM results30 seem to contradict this theory. Finally, whether the cells orientate to the microtextured surfaces as a result of the force distribution that is created by the texture of these surfaces is impossible to determine since no (known) studies have been performed to map the force distribution within cells cultured on microgrooved substratum surfaces. Recognizing the fact that these three hypotheses can even be integrated into one overall model contributes to the intriguing phenomena of cellular behavior on microtextured surfaces.

Future Perspectives: Vascular Grafts and Microtextured Surfaces As mentioned before, many investigators have already speculated on the benefits of microtextured implants. For example, Ratner1 speculated on the benefits of an implant surface that would not cause the formation of a fibrous capsule. According to Ratner, such an implant would not “be walled off ”, but the cells contacting the implant surface would respond “as if they are not seeing and interacting with the biomaterial”. This wound healing reaction would be preferable for the clinical success of several frequently used implants. For example, reduction of the capsule thickness around an implant would mean a reduction of the capsule contraction that is often observed with, for example, silicone breast implants.66 Furthermore, capsule reduction would enhance the performance of many implanted biosensors, pacemakers, and infusion pumps.17 These devices all benefit from an optimal contact between the tissues and the implant for the transduction of signals. For instance, the necessary electrical pulse of a pacemaker would be better conducted to the heart muscle if the capsule around the electrode of this device were minimized. Another good example is the sensor of an implanted insulin infusion pump. For optimal detection of insulin levels, maximal contact between the sensor of this device and the surrounding tissues is required. However, if a fibrous capsule shields the sensor from the crucial signal, the performance of the implant will be insufficient. Additional applications of microtextured biomaterials have been reported in the discipline of tissue engineering. For example, microtextured surfaces have been used in in vitro experiments to decrease hepatocyte dedifferentiation67-68 or to induce guided nerve regeneration.42,69-70 Skin autografts already have been generated out of individual keratinocytes by using orienting scaffolds,71 while speculations are voiced that guided tissue regeneration could perhaps reduce the formation of scarring tissue and enhance the repair of highly orientated structures like tendons.46,72-73 Furthermore, attempts have been made to produce large tubular morphologies with the help of tissue engineering that could function as intestine or ureter segmental replacements.74 In addition, many publications have reported on efforts to produce blood vessels.1,3,75-82 Microtextured biomaterials could be very useful as scaffolds in this field of tissue engineering research.

Microgroove Driven Tissue Ingrowth

Figure 48.12 shows the basic anatomy of a blood vessel. Large blood vessels have a thick, tough wall of connective tissue and smooth muscle, which is lined by an exceedingly thin, single layer of endothelial cells, separated from the surrounding outer layers by a basal lamina. The thickness of the connective tissue component of the vessel wall varies with the vessel’s diameter and function, but the endothelial lining is always present.83 Study of embryos has revealed that large blood vessels have all developed from small simple vessels constructed solely out of endothelial cells and a basal lamina.83 Therefore, the endothelial cells are often referred to as the pioneers, preceding the development of connective tissue and smooth muscle around the vessels when required. This is why early studies have concentrated on the formation of capillary tubes out of colonies of endothelial cells.76 However, the tissue engineering of a functional, well developed substitute blood vessel does not depend solely on the presence of vascular cells and extracellular matrix (ECM) components. Several cellular signals are equally important to processes such as cell growth and cellular differentiation. Ziegler and Nerem79 identified these signals as originating from three sources, i.e.: 1. Chemical signals, derived from the fluid (blood) flowing through the vessel; 2. Signals associated with the ECM—the ECM proteins not only hold the vascular wall together, but also participate in regulating the biology of the vascular wall; 3. The mechanical environment of the vascular wall, imposed by the hemodynamics of the vascular system. If we speculate on the roles that microtextured scaffolds could play in meeting these demands, one is evident. The orientation of the endothelial cells in the direction of the flow, and circularly orientated smooth muscle cells, could

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possibly be obtained by applying longitudinal or circular groove patterns. Not only would the overall orientation of the cell be influenced, but also the orientation of the cytoskeleton, thus contributing to the mechanical stiffness of the cells.84 By using microgrooved scaffolds, endothelial orientation could be achieved without exposing the cells to flow and cyclic stretch, phenomena that have a comparable effect on the orientation of endothelial cells.85-86 This comparison also concerns the level of gene expression. It has been reported that flow and cyclic stretch can influence the regulation of messenger RNA,87-88 an effect on the cellular metabolism that is also supposed for cells on microtextured surfaces. What kind of effect the combination of the effects of factors like surface texture, flow, and cyclic stretch will have remains to be seen, since no study investigating this combination has been performed up to this date. Another interesting option concerns the extracellular matrix proteins. The endothelial cells and smooth muscle cells in vivo are surrounded by an intricate mixture of proteins like collagens, elastin, laminin, fibronectin, and glycoaminoglycans.79 Several studies have shown that these proteins can affect the growth, differentiation, and cholesterol metabolism of both the endothelial and smooth muscle cells. Endothelial derived ECM can affect smooth muscle cell growth, depending on its composition.89 Collagen and fibronectin containing matrices promote the growth of smooth muscle cells, while matrices containing heparan sulfate proteoglycans selectively inhibit identical smooth muscle cell populations.89 Furthermore, ECM proteins can change the response of smooth muscle cells to low density lipoproteins (LDL). Earlier study90 has shown that collagen type I, for example, induces a decrease in smooth muscle cell growth, while endothelial derived ECM induces an increase in smooth muscle cell growth following incubation with

Fig. 48.12. Schematic drawing of the anatomy of a blood vessel. The longitudinal orientation of the endothelial lining (1) and the circular orientation of the smooth muscle cells (2) can be seen clearly.

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LDL.90 The ability of microtextured surfaces to enhance the production of specific proteins as mentioned earlier8 could, if true, prove a powerful tool in the creation of blood vessels through tissue engineering. These few examples show that microtextured surfaces could contribute in the study of, and the creation of, vascular grafts. This ranges from in vitro study of smooth muscle cells, where these surfaces could stop/slow down the change from a nongrowing, contractile phenotype to a proliferating, protein secreting mode,91-92 to the induction of endothelial and smooth muscle cell orientation in artificial grafts.1,75 Therefore, it is recommended that research of microtextured surfaces be exploited further in order to develop “smart” manipulative implants with higher clinical success rates. In addition, these new studies could also contribute to a better understanding of the mechanisms that cause cellular contact guidance and guidance related processes. Such insight would not only enlarge the general knowledge of these processes, but also offer implant designers a tool for designing a wider variety of more successful implants with predicable qualities. Acknowledgments This study is supported by the Technology Foundation (STW). References 1. Ratner BD. New ideas in biomaterial science—a path to engineered biomaterials. J Biomed Mater Res 1993; 27:837-850. 2. von Recum AF, Park JB. Permanent percutaneous devices. CRC Crit Rev Bioeng 1981; 5:37-77. 3. Hubbell JA. Biomaterials in tissue engineering. Biotechnology 1995; 13:565-576. 4. Langer R, Vacanti JP. Tissue engineering. Science 1993; 260: 920-926. 5. Peppas NA, Langer R. Challenges in biomaterials. Science 1994; 263:1715-172. 6. Williams DF. Definitions in Biomedicals. Progress in Biomedical Engineering. New York: Elsevier, 1987:Vol. 4. 7. Hockberger PE, Lom B, Soekarno A, Thomas CH, Healy KE. Cellular engineering: Control of cell-substrate interactions. In: Hoch HC, Jelinski LW, Craighead HG, eds. Nanofabrication and Biosystems. Cambridge: Cambridge University Press, 1996:276-299. 8. Brunette DM. Effects of surface topography of implant materials on cell behavior in vitro and in vivo. In: Hoch HC, Jelinski LW, and Craighead HG, eds. Nanofabrication and Biosystems. Cambridge: Cambridge University Press, 1996: 335-355. 9. Harrison RG. The reaction of embryonic cells to solid structures. J Exp Zool 1914; 17:521-544. 10. Loeb L, Fleisher MS. On the factors which determine the movements of tissues in culture media. J Med Res 1917; 37:75-99. 11. Weiss P. Experiments on cell and axon orientation in vitro: The role of colloidal exudates in tissue organization. J Exp Zool 1945; 100:353-386. 12. Rovensky YA, Slavnaya IL, Vasiliev JM. Behavior of fibroblast-like cells on grooved surfaces. Exp Cell Res 1971; 65:193-201.

Tissue Engineering of Prosthetic Vascular Grafts 13. Rovensky YA, Slavnaya IL. Spreading of fibroblast-like cells on grooved surfaces. Exp Cell Res 1974; 84:199-206. 14. Maroudas NG. Anchorage dependence: Correlation between amount of growth and diameter of bead, for single cells grown on individual glass beads. Exp Cell Res 1972; 74:337-342. 15. Maroudas NG. Growth of fibroblasts on linear and planar anchorages of limiting dimensions. Exp Cell Res 1973; 81:104-110. 16. Singhvi R, Stephanopoulos G, Wang DIC. Review: Effects of substratum morphology on cell physiology. Biotechnology and Bioengineering, 1994; 43:764-771. 17. von Recum AF, van Kooten TG. The influence of microtopography on cellular response and the implications for silicone implants. J Biomater Sci Polymer Edn 1995; 7:181-198. 18. den Braber ET, de Ruijter JE, Smits HTJ et al. Quantitative analysis of cell proliferation and orientation on substrata with uniform parallel surface micro grooves. Biomaterials 1996; 17:1093-1099. 19. den Braber ET, de Ruijter JE, Ginsel LA et al. Quantitative analysis of fibroblast morphology on microgrooved surfaces with various groove and ridge dimensions. Biomaterials 1996; 17:2037-2044. 20. Schmidt JA, von Recum AF. Texturing of polymer surfaces at the cellular level. Biomaterials 1991; 12:385-389. 21. Schmidt JA, von Recum AF. Surface characterization of microtextured silicone. Biomaterials 1992; 13:675-681. 22. den Braber ET, de Ruijter JE, Ginsel LA et al. Confocal laser scanning microscopical study of the fibroblast cytoskeleton, attachment complexes, and ECM protein deposition on silicone microgrooved surfaces. J Biomed Mater Res 1997; in press. 23. Benori R, Salomon D, Geiger B. Identification of two distinct domains on vinculin involved in its association with focal contacts. J Cell Biol 1989; 108:2383-2393. 24. Garbarsch C, Matthiessen ME, Olsen BE et al. Immunohistochemistry of the intercellular matrix components and the epithelio-mesenchymal junction of the tooth germ. Histochem J 1994; 26:110-118. 25. Hayman EG, Oldberg A, Martin GR et al. Co-distribution of heparan sulfate proteoglycan, laminin, and fibronectin in extra cellular matrix of normal rat kidney cells and their coordinate absence in normal cells. J Cell Biol 1982; 94:28-35. 26. Hayman EG, Pierschbacher MD, Suzuki S et al. Vitronectin; a major cell attachment-promoting protein in fetal bovine serum. Exp Cell Res 1985; 160:245-258. 27. Hayman EG, Pierschbacher MD, Ohgren Y et al. Serumspreading factor (vitronectin) is present at the cell surface and in tissues. Proc Natl Acad Sci 1983; 80:4003-4007. 28. Cox G. Trends in confocal microscopy. Micron, 1993; 24:237-247. 29. den Braber ET, de Ruijter JE, Croes HJE et al. Transmission electron microscopical study of fibroblast attachment to microtextured silicone rubber surfaces. Cells and Materials 1997; in press. 30. den Braber ET, Jansen HV, de Boer MJ et al. SEM, TEM, and CLSM observation of fibroblasts cultured on microgrooved surfaces of bulk titanium susbtrata. J Biomed Mater Res 1997; submitted. 31. den Braber ET, de Ruijter JE, Smits HTJ et al. Effect of parallel surface micro grooves and surface energy on cell growth. J Biomed Mater Res 1995; 29:511-518.

Microgroove Driven Tissue Ingrowth 32. Green AM, Jansen JA, and von Recum AF. The fibroblast response to microtextured silicone surfaces: Texture orientation into or out of the surface. J Biomed Mater Res 1994; 28:647-653. 33. Ricci JL, Charvet J, Chang R et al. In vitro effects of surface microgeometry on colony formation by fibroblasts and bone cells. 20th Annual Meeting of the Society for Biomaterials. Boston, USA. April 5-9, 1994:401. 34. Chou LS, Firth JD, Uitto VJ et al. Substratum surface topography alters cell shape and regulates fibronectin mRNA level, mRNA stability, secretion and assembly in human fibroblasts. J Cell Sci 1995; 108:1563-1573. 35. Wójciak-Stothard B, Madeja Z, Korohoda W et al. Activation of marcophage-like cells by multiple grooved substrata. Topographical control of cell behavior. Cell Biology International, 1995; 19:485-490. 36. Hong HL and Brunette DM. Effect of cell shape on proteinase secretion. J Cell Sci 1987; 87:259-267. 37. Curtis ASG, Clark P. The effects of topographic and mechanical properties of materials on cell behavior. Critical Reviews in Biocompatibility 1990; 5:344-362. 38. Chehroudi B, Gould TRL, Brunette DM. Effects of a grooved epoxy substratum on epithelial cell behavior in vitro and in vivo. J Biomed Mater Res 1988; 22:459-473. 39. Chehroudi B, Gould TRL, Brunette DM. A light and electron microscope study of the effects of surface topography on the behavior of cells attached to titanium-coated percutaneous implants. J Biomed Mater Res 1991; 25:387-405. 40. Campbell CE, von Recum AF. Microtopography and soft tissue response J Invest Surg 1989; 2:51-74. 41. Picha GJ, Drake RF. Pillared-surface microstructure and soft-tissue implants: Effect of implant site and fixation. J Biomed Mater Res 1996; 30:305-312. 42. den Braber ET, de Ruijter JE, Jansen JA. The effect of a subcutaneous silicone rubber implant with shallow surface micro grooves on the surrounding tissues in rabbits. J Biomed Mater Res 1996; in press. 42a. Clark P, Connolly P, Curtis ASG et al. Cell guidance by ultrafine topography in vitro. J Cell Sci 1991; 99:73-77. 43. Squier CA, Collins P. The relationship between soft tissue attachment, epithelial downgrowth and surface porosity. J Perio Res 1981; 16:434-440. 44. Chehroudi B., Gould TRL., and Brunette DM. The role of connective tissue in inhibiting epithelial downgrowth on titanium-coated percutaneous devices. J. Biomed. Mater. Res., 1992; 26: 493-515. 46. Ehrlich HP. Regulation der Wundheilung aus der Sicht des Bindesgewebes. Der Chirurg, 1995; 66: 165-173. 47. Baier RE. Surface properties influencing biological adhesion. Adhesion in biological systems, RS Manly (ed.) Academic Press, New York, 1970, 15-48. 48. Schakenraad JM, Busscher HJ, Wildevuur CRH et al. The influence of substratum surface free energy on growth and spreading of human fibroblasts in the presence and absence of serum proteins. J Biomed Mater Res 1986; 20:773-784. 49. Altankov G, Groth TH. Reorganization of substratumbound fibronectin on hydrophillic and hydrophobic materials is related to biocompatibility. J Mater Sci 1994; 5:732-737. 50. Williams RL, Williams DF. The spatial resolution of protein adsorption on surfaces of heterogeneous metallic biomaterials. J Biomed Mater Res 1989; 23:339-350.

543 51. Rudee ML, Price TM. The initial stages of adsorption of plasma derived proteins on artificial surfaces in a controlled flow environment J Biomed Mater Res 1985; 19:57-66. 52. Uyen HMW, Schakenraad JM, Sjollema J et al. Amount and surface structure of albumin adsorbed to solid substrata with different wettabilities in a parallel plate flow cell. J Biomed Mater Res 1990; 24:1599-1614. 53. Rapoza RJ, Horbett TA. Postadsorptive transitions in fibrinogen: Influence of polymer properties. J Biomed Mater Res 1990; 24:1263-1287. 54. Shiba E, Lindon JN, Kushner L et al. Antibody detectable changes in fibrinogen adsorption affecting platelet activation on polymer surfaces. Am J Physiol 1991; 260:C965-974. 55. Ohara PT, Buck RC. Contact guidance in vitro. Exp Cell Res 1979; 121:235-249. 56. Dunn GA, Brown AF. Alignment of fibroblasts on grooved surfaces described by a simple geometric transformation. J Cell Sci 1986; 83:313-340. 57. Izzard CS, Lochner LR. Cell-to-substrate contacts in living fibroblasts: An interference reflection study with an evaluation of the technique. J Cell Sci 1976; 21:129-159. 58. Ingber DE. Cellular tensegrity; defining new rules of biological design that govern the cytoskeleton. J Cell Sci 1993; 104:613-927. 59. Ward MD, Hammer DA. A theoretical analysis for the effect of focal contact formation on cell-substrate attachment strength. Biophys J 1993; 64:936-959. 60. Wang N, Butler JP, Ingber DE. Mechanotransduction across the cell surface and through the cytoskeleton. Science, 1993; 260:1124-1127. 61. Oakley C, Brunette DM. The sequence of alignment of microtubules, focal contacts and actin filaments in fibroblasts spreading on smooth and grooved titanium substrata. J Cell Sci 1993; 106:343-354. 62. O’Neill C, Jordan P, Riddle P et al. Narrow linear strips of adhesive substratum are powerful inducers of both growth and total focal contact area. J Cell Sci 1990; 95:577-586. 63. Oakley C, Brunette DM. Topographic compensation: Guidance and directed locomotion of fibroblasts on grooved micromachined substrata in the absence of microtubules. Cell Motility and the Cytoskeleton 1995; 31:45-58. 64. Ben-ze’ev A. The role of changes in cell shape and contacts in the regulation of cytoskeleton expression during differentiation. J Cell Sci Suppl 1987; 8:293-312. 65. Dunn GA, Heath JP. A new hypothesis of contact guidance in tissue cells. Exp Cell Res 1976; 101:1-14. 66. Bern S, Burd A, May J Jr. The biophysical and histologic properties of capsules formed by smooth and textured silicone implants in the rabbit. Plast Reconstr Surg 1992; 89:1037-1042. 67. Cima LG, Ingber DE, Vacanti JP et al. Hepatocyte culture on biodegradable polymeric substrates. Biotech. Bioengineering 1991; 38:145-158. 68. Dunn JCY, Thomkins RG, Yarmush ML. Hepatocytes in collagen sandwich: Evidence for transcriptional and translational regulation. J Cell Biol 1992; 116:1043-1053. 69. Guenard V, Kleitman N, Morrissey TK et al. Syngeneic Schwann cells derived from adult nerves seeded in semipermeable guidance channels enhance the peripheral nerve generation. J Neurosci 1992; 12:3310-3320.

544 70. Clark P, Connolly P, Curtis ASG et al. Topographical control of cell behavior: II. Multiple grooved substrata. Development 1990; 108:635-644. 71. Bell E, Rosenberg M, Kemp P et al. Recipes for reconstituting skin. J Biomech Eng Trans ASME, 1991; 113:113-119. 72. Chehroudi B, Gould TRL, Brunette DM. A light and electron microscope study of the effects of surface topography on the behavior of cells attached to titanium-coated percutaneous implants. J Biomed Mater Res 1991; 25:387-405. 73. Wòjciak B, Crossan J, Curtis ASG. Grooved substrata facilitate in vitro healing of completely divided flexor tendons. J Mat Sci Mat in Med 1995; 6:266-271. 74. Mooney DJ, Organ G, Vacanti JP et al. Design and fabrication of biodegradable polymer devices to engineer tubular tissues. Cell Transplant 1994; 3:203-210. 75. Spargo BJ, Testoff MA, Nielson TB et al. Spatially controlled adhesion, spreading, and differentiation of endothelial cells on self-assembled molecular monolayers. Proc Nat Acad Sci USA, 1994; 91:11070-11074. 76. Folkman J, Haudenschild C. Angiogenesis in vitro. Nature 1980; 288:551-556. 77. Weinberg CB, Bell E. A blood vessel model constructed from collagen and cultured vascular cells. Science 1986; 231:397-400. 78. Leenslag JW, Kroes MT, Pennings AJ et al. A compliant, biodegradable vascular graft: Basic aspects of its construction and biological performance. New Polymeric Mater 1988; 1:111- 126. 79. Ziegler T, Nerem RM. Tissue engineering of a blood vessel: Regulation of vascular biology by mechanical stresses. J Cell Biochem 1994; 56:204-209. 80. Nerem RM, Sambanis A. Tissue engineering: From biology to biological substitutes. Tissue Engineering 1995; 1:3-12.

Tissue Engineering of Prosthetic Vascular Grafts 81. Langer R, Vacanti JP, Vacanti CA et al. Tissue engineering: Biomedical applications. Tissue Engineering 1995; 1:151-161. 82. Bell E. Strategy for the selection of scaffolds for tissue engineering. Tissue Engineering 1995; 1:163-179. 83. Alberts B, Bray D, Lewis J et al. Molecular biology of the cell, 3rd edition. New York: Garland Publishing Inc., 1994. 84. Nerem RM, Girard PR. Hemodynamic influences on vascular endothelial biology. Toxicol Pathol 1990; 4:572-582. 85. Nerem RM, Levesque MJ, Cornhill JF. Vascular endothelial morphology as an indicator of blood flow. ASME J Biomech Eng 1981; 103:172-176. 86. Levesque MJ, Liepsch D, Moravec S et al. Correlation of endothelial cell shape and wall shear stress in a stenosed dog aorta. Arteriosclerosis, 1986; 6:220-229. 87. Diamond SL, Sharefkin JB, Dieffenbach C et al. Tissue plasminogen activator messenger RNA levels increase in cultured human endothelial cells exposed to laminar shear stress. J Cell Physiol 1990; 143:364-371. 88. Mitsumata M, Fishel RS, Nerem RM et al. Fluid shear stress stimulates platelet-derived growth factor expression in endothelial cells. Am J Physiol Heart Circ Physiol 1993; 265:H3-H8. 89. Hermann IM. Endothelial cell matrices modulate smooth muscle cell growth, contractile phenotype and sensitivity to heparin. Heamostasis, 1990; 20:166-177. 90. Harris-Hooker S, Sanford GL, Montgomery V et al. Influence of low density lipoproteins on vascular smooth muscle cell growth and motility: Modulation by extracellular matrix. Cell Biol Int Rep 1992; 16:433-450. 91. Chamley-Campbell JH, Campbell GR. What controls smooth muscle phenotype. Arthereosclerosis, 1981; 40:347-357. 92. Campbell JH, Campbell GR. Endothelial cell influences on smooth muscle phenotype. Annu Rev Physiol 1986; 48:295-306.

Scaffold Engineering Surface Modification

CHAPTER 49 Surface Bonding of Heparin Patrick T. Cahalan n the early 1960s Hufnagel began experiments to form “autogenized” vascular prostheses.1 This was accomplished by implanting a Teflon rod containing a loosely woven Dacron or polypropylene cloth surrounding it, and allowing 6-8 weeks for the formation of a collagen tube. Early experiments involved implantation back into the same animal to avoid immunogenic response. At the start of his work, important knowledge concerning glutaraldehyde fixation of tissue was awaiting further publications.2 “Autogenized” was replaced with the term “xenograft” due to parallel efforts by Sparks.3 Hufnagel stayed with his studies over the next two decades and moved on to attempt heparin incorporation into mandril formed prostheses as well as commercially available glutaraldehyde-treated human umbilical vein. Rigorous experimental efforts, including radioactively tagged heparin, allowed him to determine concentration of heparin in solution to effect maximal heparin loading conditions, and to his surprise he found that continued washing with saline for two hours post “imbibing” into the prosthesis did not significantly diminish the retained heparin, which was about 300 mg/ml. In implant studies the rate of early occlusion was significantly reduced, and Hufnagel reported that the half life of the heparin was limited to a few days. While the goal was to prevent thrombotic occlusion, in retrospect the most interesting observation is the ability of heparin to bind and release from collagen in an unpredicted manner. The hypothetical paradigms of the time with respect to interactions of biomolecules and surfaces were still rather narrowly focused in simple ionic charge relationships.4 Proteins (collagen) and polysaccharides (heparin) were widely known in the 70s by carbohydrate chemists to possess unique complexing capabilities that could be used to form viscosifying agents, stabilizers, and precipitation agents in food applications. The synergistic effect of these biomolecules could not be explained by simple ionic charge relationships. The tertiary and quaternary structures of biomolecules play an important role in their interactions with other molecules. While this concept is well known to anyone with an introductory education in biochemistry, it appears to be often forgotten in strategies for immobilizing biomolecules on surfaces. There remain very active efforts at producing a synthetic ionic analog heparinoid surface,5,6 while at the same time studies are showing how critical is the dependence of heparin’s bioactivity on its unique structure.7 Early attempts at heparinizing surfaces used fatty quaternary ammonium compounds as a base coating on polymers in order to present the positive charge of the quaternary amine for ionic coupling of heparin.8 These early studies clearly showed an impact of heparinized surfaces, but the heparin released rather quickly and was material dependent, suggesting that the quaternary compound (generally considered toxic) also was releasing. In attempts to make the heparin more stable, or release more slowly at the surface, different solvent systems were use to attack the base material and entangle the fatty chains with the polymer surface. Attempts were made to calculate the sustained release rate required to

I

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546

maintain thromboresistant surfaces.9 This clearly showed that it was impractical to sustain thromboresistance, due to the large loading required to maintain effectiveness for more than a few days. At the same time, Larm published his results for “end point” attached heparin on surfaces.10 This was the first time that an immobilized heparin, with no leaching or releasing heparin, could demonstrate significant deactivation of thrombin. The premise of Larm’s surface was that nitrous acid degraded heparin (NAD-Hep) containing high affinity pentasaccharides for ATIII was covalently immobilized on the surface by a terminal aldehyde group, thus presenting an optimal conformation of the molecule to the blood interface. The clearest evidence of the effectiveness of Larm’s premise is seen in controlled experiments using human blood in simulated cardiopulmonary bypass circuits where improvement in platelet protection, platelet release products, hemolysis, less white cell activation, and less contact activation is demonstrated.11 The group of Wendel has tested several coatings in the latter system, and nothing to date has been reported equal in performance to the covalently attached heparin surface. The method of Larm, commercially known as the Carmeda biologically active surface (CBAS), is in fact a covalent attachment of heparin to an adsorbed base, and thus is not completely covalent to the base material. Since Larm’s invention several additional methods of surface modification have been used to couple heparin and reported in the literature.12-18 The traditional model for the mode of action of heparin on a surface is depicted by a heparin molecule with a high affinity pentasaccharide tethered to a surface and free to bind circulating ATIII, which it conformationally alters to make it catalytic in its ability to bind and deactivate circulating thrombin. Once this action has occurred, the thrombin-antithrombin complex (TAT) is released, and the site is capable of repeating the activity. The relatively high circulating concentrations of ATIII assists in the displacement of the TAT complex from the surface.19 The question of protein adsorption and its effect on the heparin activity is a common one. Van Delden showed that the spacer composition was important for the surface retaining the ability to deactivate thrombin after exposure to plasma.20 In his studies he showed that heparin immobilized via albumin-heparin conjugates retain 27% of initial activity after exposure to blood, whereas heparin immobilized via PEO spacers can retain as much as 55% of the initial activity. Most interesting from these experiments was the fact that the total amount of protein adsorbed onto heparinized surfaces did not correlate with the retention of heparin activity. It has been reported that albumin and fibrinogen adsorbed onto heparinized surfaces are not conformationally altered.21 Earlier studies by Sevastianov suggest that it is critical to properly attach heparin if the proteins adsorbed are to be non platelet adhering and activating.22 During a panel discussion at the Biomaterials conference in New Orleans, May, 1997, a comment was made to the effect that after all these years of research on blood-compatible surfaces we are no farther than the heparinized surface. Another perspective might suggest that we have finally reached a point where we understand how to properly attach heparin to a surface. The

Tissue Engineering of Prosthetic Vascular Grafts

lessons learned should be useful insights into attachment of other biomolecules, as cited by Hubbell in attachment of peptides to surfaces.23 Our group has been studying the performance of heparinized surfaces together with the University of Maastricht, the Netherlands in attempts to understand the role of coupling methodology in the performance of these surfaces in human blood using numerous blood testing systems.24 Using a rotating disc apparatus as seen in Figure 49.1, it became evident that deactivation of thrombin at a heparinized surface is a diffusion limited reaction. The reactor was designed with baffles to assure redirection of the flow perpendicular to the test surface, which was rotating, and thus assured a uniform shear rate profile on the sample surface. A mathematical model can be hypothesized for the thrombin decay if the rate of deactivation is dependent on the diffusion to the surface, and if the concentrations of thrombin and ATIII added to the reactor, and the speed of the motor, are known. The control line in the graph of Figure 49.1 represents a blank control surface, and the limited decay of thrombin is explained by simple noncatalytic deactivation of thrombin by ATIII in solution. The heparinized surface agrees fairly well with the theoretical decay rate for a diffusion limited reaction. All surfaces that were heparinized were tested before the experiment was started to assure that no heparin was releasing. Our method of immobilizing heparin consisted of covalent attachment of hydrogels that were derivatized with amine functionality to which aldehyde-functionalized heparin (NAD-Hep, or periodate oxidized) was coupled. By controlling the amount of amine functionality we were able to prepare surfaces with different amounts of heparin coupled. Regardless of how much heparin was coupled per cm2 on the surface, or the measured bioactivity the surface possessed to deactivate thrombin, the observed behavior in Figure 49.1 was only seen at low concentrations of ATIII. At physiological concentrations of ATIII (1 U/ml), there was no difference in thrombin decay between heparinized and uncoated controls. The rate of diffusion to, and deactivation at, the surface was not as great as the noncatalyzed deactivation in solution with physiological concentrations of ATIII, even with the most active surfaces. In addition to the latter results, ELISA measurements of TAT and F1+2 fragments from blood loop experiments showed that very few of the latter complexes were formed after exposure of the heparinized surface to human whole blood in 90 minutes, as compared to uncoated control polymers. These results question the model proposed wherein thrombin in flowing blood is deactivated by the heparin-ATIII surface to form TAT complexes that are released, and the heparin repeats this activity several times. Further studies in flowing systems developed to study thromboresistance of intravascular stents, where platelet free and platelet rich plasma were used, helped solidify a refinement of the model (Figs. 49.2, 49.3, and 49.4). Figure 49.2 represents the flow through system developed for monitoring thrombin generation for an intravascular stent. Citrated PRP or PFP could be pumped through the system at controlled flow rates and samples collected over time.

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547

Fig. 49.1. Rotating disc apparatus; thrombin decay in rotating disc test.

Fig. 49.2. Flow through system which monitors Wiktor intravascular stent thrombin generation. Citrated platelet rich plasma (PRP) or platelet free plasma (PFP) is pumped through the system at controlled flow rates and samples collected over time.

548

Tissue Engineering of Prosthetic Vascular Grafts

Fig. 49.3. Comparison of contact activation as measured by factor IXa, and of thrombin generation in platelet rich plasma (PRP) or platelet free plasma (PFP). (A) Heparin-coated intravascular stents; (B) Uncoated stents.

Figures 49.3A and 49.B represent a comparison between heparin coated and uncoated intravascular stents with respect to contact activation as measured with Factor IXa, and its comparison to subsequent thrombin generation in PRP or PFP. The black points represent PRP and the gray represent PFP. The first thing that can be seen is that contact activation is independent of platelets, and that without platelets the even more “thrombogenic” metallic surface does not produce significant amounts of thrombin (Fig. 49.3A). Not surprisingly, the anionic heparin stent does cause some contact activation, but less than the noncoated stent (Fig. 49.3B). Clearly, in the presence of platelets the heparinized surface delays the onset of thrombin generation, and at equilibrium is much less thrombogenic than the noncoated control material.

Table 49.1 helps to further explain the importance of the quality (ability to bind ATIII) of the attached heparin to the surface. ATIII uptake to the surface directly correlates with the delay in the onset of thrombin generation, the amount produced at equilibrium and, most importantly, the number of adhered platelets. The platelets on the surface are determined by the LDH method, and therefore some platelets could be present that are “adhered”, but not activated. SEM photos of the surfaces of the heparinized materials did not reveal any visible platelets after mild rinsing, whereas the controls were covered with adhered and spread platelets. Much might be learned in the future by study of platelet adhesion to surfaces such as heparin and collagen. Exposed subendothelial layers of collagen and GAGs can be seen to adhere platelets in dense single cell layers that do not

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549

Fig. 49.4. Contribution of immobilized heparin to clot prevention at blood-material interface by antithrombin (ATIII) deactivation of platelet-produced thrombin.

Table 49.1. Relation between the AT uptake and the Thrombogenicity of Wiktor stents Coating

AT-Uptake

Thrombin stead State

Thrombin onset

Platelet Adhesion

None HepamedTM HepamedTM HepamedTM HepamedTM

pmol/cm-2 0.26±0.36 1.26±0.36 1.99±0.57 7.28±0.21 10.89±0.21

nM 8.83±0.37 4.98±0.62 3.40±0.25 1.75±0.09 1.48±0.15

Min 5 20 20 30 34

287±34 135±18 197±45 42±20 92±33

Table 49.2 Effects of heparin bonded surface on blood clotting in polyurethane tubing Surface

PU55D 6882-78-1 6882-78-2

Surface modification

none PU Collagen PU collagen+ heparin

Platelet Adhesion

ATIII Adsorption

Clotting Time

Plts/cm2 x103 (LDH) 123.1 234.1 195.4

pmol/cm2

seconds

nd 4.45 11.16

659 594 1183

550

appear to be spead or fully activated. This phenomenon has led us to further experiments that will be described shortly, but for now suffice it to say that the natural process of wound healing of damaged endothelium begs the question as to whether or not a natural material such as collagen in the presence of certain glycosaminoglycans might produce a “controlled” and important thrombogenic response. Thrombin generation and subsequent fibrin formation is due in a large part to the inability of flow conditions to keep the local concentration of thrombin low. Our experience has been that rapid clot formation takes place when the local concentration of thrombin reaches about 2-5 nM. Once this point has been reached, the fibrin network is rapidly formed and can entrap platelets and amplify the local generation of thrombin. Figure 49.4 is a recent attempt at refining the conceptual model of how heparin works when immobilized on a surface. The critical mechanism for preventing clotting is that thrombin produced by platelets at the surface is immediately deactivated, thus keeping the local concentration low, which in turn minimizes the strongest agent (thrombin) in platelet activation. When the thrombin level is low, it cannot amplify its own production by activating vicinal platelets, and less fibrin will minimize entrapment of additional platelets. The latter is of great significance for the longer term maintenance of low thrombogenicity, as thrombin can adhere to fibrin and in this state it is not deactivated by heparin and ATIII, and can continue to activate platelets.25 Intramuscular implant studies of immobilized collagen on biomaterials by our group showed immediate thrombogenic response followed by strong apparent neutrophil recruitment at the tissue-material interface. This response quickly became quiescent, and after 6 weeks the tissue-material interface had only an occasional macrophage with mostly fibroblasts (two cell layers) followed by adjacent microvascular structures. The control material (polyurethane) had a typical foreign body fibrotic response with macrophages and foreign body giant cells adjacent to the majority of the surface, at least 15 layers of densely packed fibroblasts with no vascular structures, and loose connective tissue outside of the fibroblast layers. These results led to preliminary experiments in measuring the coagulation response to collagen immobilized surfaces, and the potentiating effect that heparin coupled to this collagen layer might have. Using the same flow through system developed at the University of Maastricht and described previously, polyurethane tubings were prepared with immobilized collagen and immobilized collagen with heparin covalently bonded to the collagen (type I collagen). Apparently the fact that more platelets are attached to the collagen and collagen plus heparin surfaces than the polyurethane does not lead to a faster clotting time, and in the case of the heparinized collagen the clotting time was significantly delayed (Table 49.2). These are very preliminary experiments, and only suggest additional studies to attempt to look at heparin in a broader context of the total healing response.

Tissue Engineering of Prosthetic Vascular Grafts

Heparin has several other properties in addition to acting as an anticoagulant, such as antibacterial and antiviral activity, inhibition of several enzymes, and the stimulation of lipoprotein lipase release from the surface of endothelial cells.26,27 It has been proposed that heparin has a role in the defense against pathogens outside of the immune system.28 This hypothesis seems to have merit based on the fact that heparin is predominantly located in organs and tissues that come in direct contact with the outside environment (lung, skin, and intestine). Heparin binding growth factors (HBGFs) represent an important family of mediators for the healing response, capable of inducing mesenchymal cell proliferation and differentiation, tissue regeneration, morphogenesis, and neovascularization.29 Several investigations of the growth factors that are found in wound sites use heparin columns in separation and identification of these growth factors.30-32 Numerous experiments have been tried by simple addition of growth factors to wound sites or to matrices. There are also some experiments where the role that heparin might play has been investigated. Angiogenesis was found to be enhanced by the addition of heparin to gels made of basement membrane extract that also had added acidic and basic fibroblast growth factors.33 Heparin, along with ECGF, is suggested to play a role as a response modifier of human endothelial cell migration, which may be relevant to tumor metastasis, wound healing, and atherogenesis.34 Another study using heparin analogs (chemically substituted dextrans) in a collagen plaster was able to show a remarkable effect both on the kinetics and on the quality of the restored skin.35 This study strongly suggests that endogenous growth factors naturally releasing during the regeneration process could be trapped, protected and released by the addition of heparin analogs or heparin. It may be wishful thinking, but perhaps immobilized or surface bonded heparin may play a strong role in tissue engineering. This may be particularly true if it is properly immobilized and capable of trapping, protecting and releasing growth factors at locations and rates desired. Realization of this potential probably will require some additional thinking beyond heparin by itself. Coupling to biomolecules such as collagens and other glycosaminoglycans to polymeric surfaces may prove fruitful. From some of the studies mentioned, it is apparent that gel matrices can be modified to exhibit angiogenesis and remodeling. It may be entirely possible to fill porous polymer matrices with such gels and approach the ultimate regenerated tissue-like vascular graft. Still, for those concerned with optimal mechanical properties, the proper polymeric scaffold with a more fibrous character can have the surfaces modified for covalent attachment that may achieve a similar end. The former will require a bigger step in regulatory and commercial acceptance, while the latter may lead to incremental improvements. It is hopeful that efforts in both areas will lead to understanding that will result in a vascular graft with complete healing that is long term patent, and infection resistant.

Surface Bonding of Heparin

References 1. Gomes MN, Hufnagel CA, Dokumaci O. Studies in the autogenization of prosthesis for samll artery replacement. Am Surg 1972; 38:664-666. 2. Russel AP, Hopwood D. The biological uses and importance of glutaraldehyde. Prog Med Chem 1976; 13:271-299. 3. Sparks CH. Silicone mandril method for growing reinforced autogenous femoropopliteal artery grafts in situ. Ann Surg 1973; 177:293-300. 4. Bothwell JW, Lord GH, Rosenberg N. Modified arterial heterografts: Relationship of processing techniques to interface characteristics. In: Sawyer PN, ed. Bio-physical Mechanisms in Vascular Homeostasis and Intravascular Thrombosis. Appleton-Century-Crofts, 1965. 5. Csomor K, Karpati E, Nagy M, Gyorgyi-Edeleny J, Machovich R. Blood coagulation is inhibited by sulphated copolymers of vinyl alcohol and acrylic acid under in vitro as well as in vivo conditions. Thromb Res 1994; 74(4):389-398. 6. Montdargent B, Toufik J, Carreno M, Labarre D, Jozefowicz M. Complement activation and adsorption of protein fragements by functionalized polymer surfaces in human serum. Biomater 1992; 13. 7. van Boeckel CAA. From Heparin To A Synthetic Drug: A Multi-disciplinary Approach. In: Trends in Receptor Research. Elsevier Publishers B.V. 1993. 8. Gott VL, Koepke DE, Daggett RL, Zarnstorff W, Young WP. The coating of intravascular plastic prostheses with colloidal graphite. Surg 1961; 50:382. 9. Bassmadijan D, Sefton MV. Relationship between release rate and surface concentration for heparinized materials. J Biomed Mat Res 1983; 17:509-518. 10. Larm O, Larsson R, Olsonn P. A new nonthrombogenic surface prepared by selective covalent binding of heparin via a modified reducing terminal residue. Biomat Med Dev Art Org 1983; 11:161-173. 11. Wendel HP, Heller W, Gallimore MJ, Hoffmeister HE. Heparin-coated oxygenators significantly reduce contact system activation in an in vitro cardiopulmonary bypass model. Blood Coag Fibrin, 1994; 5. 12. Grainger DW, Kim SW, Feijen J. Poly(dimethylsiloxane)poly(ethylene oxide)-heparin block copolymers. I. Synthesis and characterization. J Biomed Mater Res 1990; 22:231-249. 13. Ebert CD, Kim SW. Immobilized heparin: spacer arm effects on biological interactions. Thromb Res 1982; 26:43-57. 14. Yuan S, Szakalas-Gratzl G, Ziats NP, Jacobsen DW, Kottke-Mrchant K, Marchant RE. Immobilization of high affinity heparin oligosacccharide to radio frequency plasma-modified polyethylene. J Biomed Mater Res 1993; 27:811-819. 15. Chandler WL, Solomon DD, Hu CB, Schmer G. Estimation of surface bound heparin activity: A comparison of methods. J Biomed Mater Res 1988; 22:497-508. 16. Engbers GHM. The development of heprinized materials with an imporved blood compatibility. Thesis University of Twente, The Netherlands, 1990. 17. Cahalan L et al. Biocompatible medical article and method. U.S. Patent 5,607,475, Assignee: Medtronic Inc. March 4, 1997.

551 18. Yun JK, DeFife K, Colton E, Stack S, Azeez A, Cahalan L, Verhoeven M, Cahalan P, Anderson JM. Human monocyte/macrophage adhesion and cytokine production on surface-modified poly(tetrafluoroethylene/hexafluoropropylene) polymers with and without protein preadsorption. J Biomed Mater Res 1995; 29:257-268. 19. Elgue G et al. Effect of surface immobilized heparin on the activation of adsorbed factor XII. Art Org 1993; 17(8):721-726. 20. van Delden CJ. On the anticoagulant activity of heparinized surfaces. Ph.D. Thesis, University of Twente, The Netherlands, 1995. 21. Barbucci R, Magnani A. Conformation of human plasma proteins at polymer surfaces: The effectiveness of surface heparinization. Biomater 1994; 15(12):955-962. 22. Nemets EA, Sevastianov V. The interaction of heparinized biomaterials with human serum, albumin, fibrinogen, atntithrombin III, and platelets. Art Org 1991; 15(5):381-385. 23. Drumheller PD, Hubbell JA. Densely crosslinked polymer networks of poly(ethylen glycol) in trimethylolpropane triacrylate for cell-adhesion-resistant surfaces. J Biomed Mater Res 1995; 29:207-215. 24. Lindhout T, Blezer R, Schoen P, Willems GM, Fouache B, Verhoeven M, Hendiks M, Cahalan L, Cahalan P. Antithrombin activity of surface-bound heparin studied under flow conditions. J Biomed Mater Res 1994; 29:1255-1266. 25. Buchanan MR, Liao P, Ofosu FA. Simultaneous inhibition of thrombin by ATIII and HCII and prevention of thrombus formation and growth: Relative effects of heparin and sulodexide. Abstract from XIV Congress of ISTH, 1993. 26. Wright TC, Castellot JJ, Diamond JR, Karnovsky JJ. Regulation of cellular proliferation by heparin and heparan sulphate, In: Lane DA, Lindahl U, eds. Heparin, chemical and bioligical properties, clinical applications. London: Edward Arnold, 1989. 27. Olivecrona T, Bengtsson-Olivecrona G. Heparin and lipases. In: Lane DA, Lindahl U, eds. Heparin, chemical and biological properties, clinical applications London: Edward Arnold, 1989. 28. Björk I, Olson ST, Shore JD. Molecular mechanisms of the accelaerating effect of heparin on the reactions between antithrombin and clotting proteinases. In: Lane DA, Lindahl U, eds. Heparin, chemical and bioligical properties, clinical applications. London: Edward Arnold, 1989. 29. Lobb RR. Clinical applications of heparin-binding growth factors. Euro J Clin Inves 1988; 18(4):321-36. 30. Hamerman D, Taylor S, Kirschenbaum I, Klagsbrun M, Raines EW, Ross R, Thomas KA. Growth factors with heparin binding affinity in human synovial fluid. Proc Soc Exp Biol Med 1987; 186(3):384-9. 31. Marikovshy M, Breuing K, Liu PY, Eriksson E, Higashiyama S, Farber P, Abrahm J, Klagsbrun M. Appearance of heparin-binding EGF-like growth factor in wound fluid as a response to injury. Proc Natl Acad Sci USA 1993; 90(9):3889-93. 32. McCarthy DW, Downing MT, Brigstock DR, Luquette MH, Brown KD, Abad MS, Besner GE. Production of heparin-binding epeidermal growth factor-like growth factor (HB-EGF) at sites of thermal injury in pediatric patients. J Inves Dermatol 1966; 106(1):49-56.

552 33. Passaniti A, Taylor RM, Pili R, Guo Y, Long PV, Haney JA, Paul RR, Grant DS, Martin GR. A simple, quantitative method for assessing angiogenesis and antiangiogenic agents using reconstituted basement membrane, heparin, and fibroblast growth factor. Lab Inves 1992; 67(4):519-28. 34. Terranova VP, DiFlorio R, Lyall RM, Hic S, Friesel R, Maciag T. Human endothelial cells are chemtactic to en-

Tissue Engineering of Prosthetic Vascular Grafts dothelial cell growth factor and heparin. J Cell Biol 1985; 101(6):2330-4. 35. Meddahi A, Blanquaert F, Saffar JL, Colombier ML, Caruelle JP, Josefonvicz J, Barritault D. New approaches to tissue regeneration and repair. Path Res Prac 1994; 190(9-10):923-8.

Scaffold Engineering Surface Modification

CHAPTER 50 Covalent Grafting of RGD Peptides to Synthetic Surfaces Nina M.K. Lamba, S.L. Cooper

Introduction

F

abric materials were first used as vascular prostheses in the 1950s, when Voorhees et al implanted a polymeric vascular graft manufactured from vinyl chloride and acrylonitrile.1 Since then, a number of polymers have been used to fabricate vascular prostheses, and today, polyethylene terephthalate (PET, Dacron) and polytetrafluoroethylene (PTFE, GoreTex) are the most commonly used biomaterials for this application. Vascular grafts made from either of these materials have an internal diameter usually greater than 6 mm, and are restricted to use in high shear environments. Currently there is no clinically acceptable synthetic small diameter vascular prosthesis, i.e., one less than 6 mm I.D. Thrombus and neointima formation will readily occlude synthetic vascular grafts. The resulting loss of patency remains the greatest obstacle to the development of a small caliber vascular prosthesis. Thus, there is a need for an engineered material that will provide the necessary mechanical and thromboresistant properties for a successful synthetic vascular prosthesis, to facilitate the advancements in vascular surgery. In this chapter we will discuss the motivation for the endothelialization of luminal surfaces, and describe some of the approaches that have been taken to achieve this. A discussion of methods to covalently graft Arg-Gly-Asp (RGD) peptides to polymers follows, focusing on methods and results obtained using polymers that are used to fabricate vascular prostheses.

Methods to Improve Endothelialization The only nonthrombogenic material known to man is the normal vascular endothelium. The endothelium actively achieves this nonthrombogenicity through the synthesis of numerous anti-platelet agents and clotting inhibitors. Furthermore, protein adsorption occurs rapidly onto artificial surfaces from solution, but this is not a feature of endothelial surfaces.2-4 There is a great deal of literature detailing various approaches to promote the development of a stable, confluent layer of endothelial cells on the inner lumen of a vascular graft. It is believed that the growth of a confluent monolayer of endothelial cells onto a synthetic substrate will not only reduce the degree of thrombus formation, but may also reduce the proliferation of pathogenic organisms on the device. Approaches to encourage the endothelialization of the inner lumen of vascular grafts have included seeding of surfaces with cells, and chemical and physical modification of surfaces to promote cell attachment. Prostheses have been seeded with endothelial cells to promote endothelialization of the luminal surface. The success of this approach has been limited by problems with harvesting endothelial cells, and the subsequent attachment, retention and proliferation of cells on the surface. Chemical modification of surfaces has included the use of photo discharge Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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technology to deposit reactive groups onto polymer surfaces, or to indirectly influence the proteins adsorbed to the surface. These methods to encourage cell attachment have met with limited success, owing to the lack of specificity and poor control over protein orientation. Physical methods to improve the development of a pseudointima on the surface have involved altering the porosity and texture of surfaces, to promote tissue ingrowth into the graft. The neointima that forms in the lumen of a graft may not provide comparable physiological properties of a natural endothelium, limiting the success of integration of the implanted device with the host by this method.5 Grafts have been coated with adhesive proteins or extracellular matrix components in order to promote attachment and proliferation of endothelial cells.6-8 An improvement in the extent of endothelialization has been observed in animal models. RGD-protein conjugates have also been synthesized,9 but problems in defining the precise conformation of the protein once it has adsorbed to the substrate impede interpretation of results. The long term success of materials modified through adsorption of proteins to mediate cellular adhesion is not guaranteed, due to possible desorption of the protein or proteolysis. More recently, there has been an interest in the grafting of synthetic peptide sequences onto polymeric substrates. By covalently grafting short peptide sequences onto polymeric substrates, it is believed that cellular adhesion can be mediated directly through receptor-ligand interactions, rather than through a conditioning layer of adsorbed protein. The covalent grafting of the Arg-Gly-Asp (RGD) peptide sequence to synthetic materials has also been shown to promote cell adhesion and attachment for applications such as tissue culture substrates and biomaterials for implantation and hybrid artificial organs. Some of the approaches to modify biomaterials with cell adhesive peptide sequences containing the RGD sequence relevant to the development of synthetic vascular grafts will be discussed in more detail in this chapter.

Immobilization of RGD Peptides The RGD peptide sequence has been shown to be the minimal cell-recognizable sequence in many adhesive plasma and extracellular matrix proteins.10 The RGD sequence was first found in fibronectin, and is also present in vitronectin, von willebrand factor, fibrinogen, and collagen. The RGD sequence has been shown to play a crucial role in mediating cell attachment and subsequent spreading. It has been shown to be adhesive towards platelets and other cells, including fibroblasts and endothelial cells. The RGD tripeptide may also act as a ligand for the integrin superfamily of receptors. In particular, many integrin receptors recognize the RGD sequence which is present in many adhesive proteins.11 It has been demonstrated that the synthetic RGD peptide in solution is able to compete with adhesive proteins adsorbed onto a surface for binding receptors on the cell in solution and prevent cellular adhesion to an adhesive protein adsorbed to the surface.10 Synthetic surfaces containing immobilized RGD peptides have been shown to promote cell attachment in a manner similar to fibronectin. Some cell surface receptors have been shown to bind with the RGD

Tissue Engineering of Prosthetic Vascular Grafts

sequence in a specific protein, whereas other receptors may recognize the RGD sequence in more than one protein. The specificity of the RGD protein is believed to be modulated by the conformation of the sequence. This may be determined by the amino acids that are immediately adjacent to the RGD sequence.12-14 Substitution of peptides within the RGD sequence has been shown to produce a large reduction in the adhesivity of the peptide.15 The presence of a peptide adjacent to the sequence has been shown to alter the activity of the RGD sequence.10,16,17 Covalent grafting of biological molecules to a synthetic substrate prevents the desorption of the agent over time from the surface. An excellent review of the chemistry of covalent immobilization of RGD peptides is available.18 Generally, either the peptide or the surface must be “activated” to provide suitable sites for immobilization of the ligand to the substrate. Biological structures are less resistant than synthetic polymers to harsh chemical environments. Prolonged exposure of a ligand to a harsh environment may lead to a reduction in the biological activity of the peptide. Thus, the polymer substrate is often derivatized or activated prior to peptide coupling. Some of the more commonly used reaction schemes involve the coupling of the peptide sequence to a polymer substrate that contains hydroxyl, thiol, carboxylic acid or amine groups. Surfaces or ligands can also be derivatized and immobilized using photochemical techniques. The efficiency of the coupling reaction between the substrate and the peptide can influence the surface peptide density, which has been shown to influence the degree of cell spreading.19,20 There are also other factors that can affect the coupling yield and biological activity of the ligand. The RGD sequence can be immobilized by covalent bonding at either end of the sequence, either through the carboxyl terminus or the amino terminus. The orientation of the sequence can affect the interactions of the peptide with cell surface receptors. The immobilization of peptide sequences on surfaces may also impose steric constraints on the peptide, affecting the affinity and specificity of the ligand. This has been reported with the grafting of other biological molecules, and may be overcome by grafting the ligand onto spacer arms, reducing steric hindrance imposed by the proximity of the ligand to a rigid surface.21 The protection of the peptide during the coupling reaction scheme is also a consideration, to ensure that the activity of the peptide is retained. Competing side reactions such as hydrolysis, polymerization and crosslinking can also affect the grafting yield. The final surface density of ligands can be determined by radiolableling, photochrome labeling, surface analysis or gravimetry.18 A number of studies have been performed investigating methods to immobilize peptide sequences containing the RGD sequence to polymeric surfaces. With respect to the development of synthetic vascular prostheses, RGD grafted materials have been studied with the goal of improving the integration of the prosthesis into the host cardiovascular system, with the ability to perform the biological functions of the endothelium. RGD sequences have been grafted onto polyethylene terephthalate (PET) and polytetrafluoroethylene (PTFE),22 polyvinyl alcohol (PVA),23 poly-

Covalent Grafting of RGD Peptides to Synthetic Surfaces

acrylamide,24 polystyrene,25 and polyurethane.26 It has been demonstrated that the presence of RGD peptide sequences at the surface increases endothelial cell attachment in vitro. The attachment of endothelial cells through specific receptor-ligand interactions should overcome the problems of cell detachment under shear conditions.22 The methods to immobilize RGD peptides to polymers, described below, demonstrate that successful grafting can be achieved by either bulk or surface modification. The main advantages of using surface modification techniques are that smaller quantities of reagents are required, and the mechanical properties of the material can be retained. However, on a commercial scale, bulk modification of a material requires fewer processing steps, which can ease fabrication and reduce costs. Sugawara and Matsuda have derivatized peptides with 4-azidobenzoyloxysuccinimide. The peptides were then adsorbed to a polyvinylalcohol surface.27 The surface was exposed to UV radiation to covalently bond the peptides to the surface. Bovine aortic endothelial cell adhesion was reported to be enhanced in a biologically specific manner. Ozeki and Matsuda have used a similar strategy to attach photochemically peptides containing the RGD sequence to polystyrene.25 An increase in the attachment of bovine aortic endothelial cells was reported. The specific peptide sequence that is grafted may also modulate the adhesivity of the RGD sequence. Hirano et al16 immobilized RDGS, RGDV, RGDT, and RGD onto ethylene-acrylic acid copolymer by coupling the terminal NH2 to the carboxylic residues in the polymer. The surfaces were evaluated for cell recognition by cell adhesion activity towards 5 different cell lines. It was found that the fourth amino acid in the tetrapeptides played an important role in determining the precise specificity of the RGD sequence. The presence of either serine (S), threonine (T), or valine (V) at the end of the peptide enhanced cellular adhesion. They postulated that the modulation in the adhesive nature of the RGD sequence may arise from a conformational change in the RGD sequence, exerted by the terminal peptide. Massia and Hubbell have grafted RGD peptides onto the surface of polyethylene terephthalate (PET) and polytetrafluoroethylene (PTFE) films.22,28 This was achieved through the hydroxylation of the polymer followed by subsequent immobilization of the synthetic peptide sequence via tresyl chloride. The RGD could only react via the N-terminus, ensuring that the orientation of the grafted peptide was the same throughout. Radiolabeling of peptides was used to confirm the presence and concentration of the peptide sequence. Endothelial cell adhesion was evaluated on the unmodified, hydroxylated and peptide containing surfaces, both with and without the presence of serum. In the absence of serum, the RGD grafted films were able to support the adhesion and spreading of human umbilical vein endothelial cells (HUVECs). Neither the unmodified films nor the hydroxylated intermediates showed the same degree of adhesion and spreading. Stress fibers were also observed, implying that the cells were attached strongly to the substrate, and may possess the mechanical strength to resist detachment by shear forces.

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Over the years, it has become apparent that the thrombogenicity of a synthetic graft is not the only criterion by which a vascular graft should be assessed. The mechanical properties, particularly the compliance and the texture of the graft are important factors in promoting tissue ingrowth, reducing anastomotic hyperplasia and favoring the long term patency of the graft. Both PET and PTFE are much less distensible when compared with the natural artery wall. Natural arteries show an increase in diameter of about 10% when pressurized to 150 mm Hg (normal arterial pressure). In comparison, PET and PTFE grafts distend by about 1% under these conditions. Polyurethanes are one of the strongest candidates for the development of small diameter vascular prostheses, due to their high tensile strength, compliance and blood compatibility.29,30 Polyurethanes have been shown to distend by about 6% under normal arterial pressure, 31 and a blended polyurethanepolylactide vascular prosthesis possesses mechanical properties similar to rat abdominal aorta.32 Arterial substitutes need to have mechanical properties that match the natural vessel wall, to avoid turbulent blood flow, which may result in thrombus formation or destruction of formed blood elements. Lyman et al33,34 have shown that the distensibility of the graft can influence the biocompatibility through the degree of anastomotic hyperplasia by comparing vascular grafts fabricated from the same material, but with different degrees of compliance. Furthermore, a mismatch in compliance between the synthetic graft and the natural vessel may traumatize the natural vessel and disrupt the endothelium, which will in turn initiate thrombosis and stimulate intimal hypertrophy. Polyurethanes are block copolymers consisting of alternating soft and hard segments. The incompatibility between these two components of the polymer allows microphase separation to occur within the bulk polymer. The material can be described as being comprised of hard domains dispersed within a soft segment matrix. This gives rise to the superior physical properties of the polyurethanes and is believed to be a contributory factor to their blood and tissue compatibility. Recently, RGD peptide sequences have been successfully grafted to polyurethane substrates.35 Polyurethanes for medical applications are usually synthesized via a two-step ‘prepolymer’ method. Lin et al synthesized a polyurethane from methylene bis (p-phenylisocyanate) (MDI) and butane diol (BD), with polytetramethyleneoxide (PTMO) as the soft segment. Hydrogen atoms on the urethane groups were substituted with carboxyl groups, via abstraction of urethane hydrogen by sodium hydride, followed by reaction of the polyurethane with β-propiolactone. The degree of substitution could be varied by the quantities of reagents used. This was performed to alter the density of the grafted peptides. Peptide sequences were then coupled to the carboxyl group using 1-(3dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDCl). The reaction scheme for the carboxylation of the urethane group, followed by coupling of the RGD peptide sequence is shown in Figure 50.1. Lin investigated the effect of the coupling reaction on the endothelialization of the substrate. Two coupling

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Tissue Engineering of Prosthetic Vascular Grafts

Fig. 50.1. Reaction scheme for the synthesis of RGD-containing peptide grafted polyurethane. (a) deprotonation of urethane group; (b) carboxylation of polyurethane; (c) coupling of RGD peptide.26

methods were used. The first, one step method coupled the free hexapeptide directly to the carboxyl group through the formation of an amide bond. The second method involved two steps, coupling a protected peptide to the carboxyl group, followed by deprotection of the peptide. The two step method using the protected peptides appeared to have a higher coupling efficiency of the peptide to the substrate of the unprotected peptide, as coupling could only occur via the N-terminus of the peptide. Furthermore, there were fewer competing side reactions in the synthetic scheme using the protected peptides, improving the yield of grafting. Previously, other researchers had reported that the adhesion and growth of endothelial cells on polyurethane is relatively poor.36,37 The immobilization of RGD peptides to a polyurethane as described above increased the number of adherent HUVECs after in vitro seeding. A larger number of cells attached to RGD peptides that had been protected during synthesis than those that had not been protected. In addition, greater endothelial cell adhesion to hydrated samples was observed compared with samples that had not been previously hydrated. Polyurethanes contain hydrophilic and hydrophobic domains that reorient upon hydration, minimizing the interfacial free energy. Cold-stage ESCA studies of hydrated and dry samples showed that reorientation of the peptides occurs at the hydrated surface,38 leading to the conclusion that the RGD peptides on the polyurethane backbone do orient towards the surface on hydration. This reorientation of the functional peptide groups is depicted schematically in Figure 50.2.38 Reorientation of RGD peptides at an aqueous interface has been reported else-

where.25 The increased adhesion on the RGD-grafted polyurethanes that were hydrated prior to contact with endothelial cells provides further evidence that reorientation of the RGD peptides does occur on hydration, and that cellular adhesion to RGD grafted materials is mediated through receptor interactions with the peptide. Figure 50.3 shows SEMs of endothelial cells grown in vitro, on the base polyurethane (PEU), carboxylated polyurethane (PEU-COOH), and peptide grafted polyurethanes (PEUGRGESY, PEU-GRGDSY, and PEU-GRGDVY). The effect of substituting one of the amino acids within the RGD sequence reduces the adhesivity of the peptide (PEUGRGESY). The substitution of the peptide adjacent to the RGD sequence has the ability to alter the adhesivity, as adhesion was greater on GRGDVY grafted material than GRGDSY grafted polyurethane. Both Massia22 and Lin35 have investigated the role of serum proteins in mediating cell adhesion to RGD grafted materials. Endothelial cell adhesion was evaluated on the unmodified, intermediate and peptide grafted materials, both with and without the presence of serum. Massia and Hubbell found that both the unmodified PET and the hydroxylated PTFE supported cell adhesion only in the presence of serum proteins, implying that cell adhesion was mediated by adsorbed proteins. Lin et al35 reached a similar conclusion in their study of HUVEC attachment to the RGD grafted polyurethane. Hubbell and co-workers have also created highly specific cell adhesive substrates of RGD-containing surfaces with low protein adsorption.15,22,39 The low protein adsorbing characteristics of these surfaces has allowed

Covalent Grafting of RGD Peptides to Synthetic Surfaces

Fig. 50.2. Schematic of the surface re-orientation of the peptide grafted polyurethanes in the hydrated and dehydrated states. From: Lin H-B, Lewis KB, Leach-Scampavia D et al. Surface properties of RGDpeptide grafted polyurethane block copolymers: Variable take-off angle and cold-stage ESCA studies. J Biomater Sci, Polym Ed 1993; 4:183-198.

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Fig. 50.3. SEMs of HUVECs attached to polyurethanes. Base polyurethane (PEU), carboxylated polyurethane (PEUCOOH), RGE-grafted polyurethane (PEU-GRGESY), RGD grafted polyurethane (PEUGRGDSY and PEU-GRGDVY). From: Lin H-B, GarciaEcheverria C, Asakura S et al. Endothelial cell adhesion on polyurethanes containing covalently attached RGD peptides. Biomaterials 1992; 13:905-914.

the study of cell attachment and spreading independently of protein adsorption. Using these substrates they have shown that the RGD and Tyr-Ile-Gly-Ser-Arg (YIGSR) peptide sequences promote endothelial cell adhesion independently of adsorbed cell adhesion proteins. They have also demonstrated that serum proteins play a role in cell spreading, with serum proteins affecting the final extent of spreading, but not the initial rate.40 The effect of the surface density of the grafted peptide has also been investigated using these surfaces. The effect of ligand surface density on cell adhesion is discussed elsewhere in this book. Ultimately, it is hoped that optimization of the spacing of RGD peptides to promote the attachment of endothelial cells will allow the inclusion of other ligands, such as growth factors, that may be necessary to further modulate adhesion and spreading, or otherwise maintain the cell layer.

Summary The natural endothelium offers the only nonthrombogenic surface known to man. This is actively achieved through regulation of transport across the endothelium, and the synthesis of agents that actively prevent events such as protein adsorption, platelet deposition and thrombus formation. The development of a successful small diameter vascular prosthesis is central to the advancement of peripheral vascular surgery. The issues of occlusion and infection of vascular grafts are problematic, and remain the greatest impediment to the development of small caliber grafts. The immobilization of RGD peptides to the luminal surfaces of synthetic vascular prostheses may offer the means to promote endothelialization, and improve the long term patency of vascular grafts. Further studies are required to establish the long term clinical performance of such devices. Studies need to include examination of endothelial cell retention once the material is exposed to a high shear envi-

ronment, the physiological properties of the endothelial layer, and the interaction of endothelial cells with the formed elements of the blood, such as platelets and white blood cells. References 1. Voorhees AB, Jaretzki A, Blakemore AH. The use of tubes constructed from Vinyon N cloth in bridging arterial defects. Ann Surg 1952; 135:322-324. 2. Szycher M. Thrombosis, hemostasis and thrombolysis at prosthetic interfaces. In: Szycher M, ed. Biocompatible Polymers, Metals and Composites. Lancaster, PA: Technomic, 1983:1-33. 3. Brash JL. Role of plasma protein adsorption in the response of blood to foreign surfaces. In: Sharma CP, Szycher M, ed. Blood compatible materials and devices. Lancaster, PA: Technomic, 1991:3-24. 4. Forbes CD, Courtney JM. Thrombosis and artificial surfaces. In: Bloom AL, Forbes CD, Thomas AL, Tuddenham EGD, eds. Haemostasis and Thrombosis. London: Churchill Livingstone, 1993:1301-1324. 5. Zilla P, Fasol R, Grimm M, Fischlein T, Eberl T, Preiss P, Krupicka O, Oppell U, Deutsch M. Growth properties of cultured human endothelial cells on differently coated artificial heart materials. J Thorac Cardiovasc Surg 1991; 101:671-680. 6. Lee Y, Park DK, Kim YB, Seo JW, Lee KB, Min B. Endothelial cell seeding onto the extracellular matrix of fibroblasts for the development of a small diameter polyurethane vessel. ASAIO J 1993; 39:M740-M745. 7. Miwa H, Matsuda T, Tani N, Kondo K, Iida F. An in vitro endothelialized compliant vascular graft minimizes anastomotic hyperplasia. ASAIO J 1993; 39:M501-M505. 8. Stansby G, Berwanger C, Shukla N, Schmidt-Rixen T, Hamilton G. Endothelial seeding of compliant polyurethane vascular graft material. Brit J Surg 1994; 81:12861289.

Covalent Grafting of RGD Peptides to Synthetic Surfaces 9. Kishida A, Takatsuka M, Matsuda T. RGD-albumin conjugate: Expression of tissue regeneration activity. Biomaterials, 1992; 13:924-930. 10. Pierschbacher MD, Ruoslahti E. Cell attachment activity of fibronectin can be duplicated by small synthetic fragments of the molecule. Nature 1984; 309:30-34. 11. Hynes RO. Integrins: A family of cell surface receptors. Cell 1987; 48:549-559. 12. Ruoslahti E, Pierschbacher MD. New perspectives in cell adhesion: RGD and integrins. Science 1987; 238:491-497. 13. D’Souza SE, Ginsberg MH, Plow EF. Arginyl-glycyl-aspartic acid (RGD): A cell adhesion motif. Trends Biol Sci 1991; 16:246-250. 14. Hautanen A, Gailit J, Mann M, Ruoslahti E. Effects of modification of the RGD sequence and its context on recognition by the fibronectin receptor. J Biol Chem 1989; 264:1437-1492. 15. Drumheller PD, Hubbell JA. Polymer networks with grafted cell adhesion peptides for highly specific cell adhesive substrates. Anal Biochem 1994; 222:380-388. 16. Hirano Y, Okuna M, Hayashi T, Goto K, Nakajima A. Cell-attachment activities of surface immobilized oligopeptides RGD, RGDS, RGDV, RGDT, and YIGSR towards five cell lines. J Biomater Sci Polym Ed 1993; 4:235-243. 17. Lin H-B, Garcia-Echeverria C, Asakura S, Sun W, Mosher DF, Cooper SL. Endothelial cell adhesion on polyurethanes containing covalently attached RGD peptides. Biomaterials 1992; 13:905-914. 18. Drumheller PD, Hubbell JA. Surface immobilization of adhesion ligands for investigations of cell-substrate interactions. In: Bronzino JD, ed. The Biomedical Engineering Handbook. Boca Raton, FL: CRC Press and IEEE Press, 1995:1583-1596. 19. Massia SP, Hubbell JA. An RGD spacing of 440 nm is sufficient for integrin αvβ3−mediated fibroblast spreading and 140 nm for focal contact and stress fiber formation. J Cell Biol 1991; 114:1089-1100. 20. Goodman SL, Cooper SL, Albrecht RM. Integrin receptors and platelet adhesion to synthetic surfaces. J Biomed Mater Res 1993; 27:683-695. 21. Nojiri C, Okano T, Park KD, Kim SW. Suppression mechanisms for thrombus formation on heparin-immobilized segmented polyurethane-ureas. Trans Am Soc Artif Intern Organs 1988; 34:386-398. 22. Massia SP, Hubbell JA. Human endothelial cell interactions with surface coupled adhesion peptides on a nonadhesive glass substrate and two polymeric biomaterials. J Biomed Mater Res 1991; 25:223-242. 23. Matsuda T, Kondo A, Makino K, Akutsu T. Development of a novel artificial matrix with cell adhesion peptide for cell culture and artificial hybrid organs. Trans Am Soc Artif Intern Organs 1989; 35:677-679. 24. Brandley BK, Schnaar RL. Covalent attachment of an ArgGly-Asp sequence peptide to derivatizable polyacrylamide surface: Support of fibroblast adhesion and long-term growth. Anal Biochem 1988; 172:270-278. 25. Ozeki E, Matsuda T. Development of an artificial extracellular matrix. Solution castable polymers with cell rec-

559 ognizable peptidyl side chain. Trans Am Soc Artif Intern Organs 1990; 36:M294-M296. 26. Lin H, Zhao Z, Garcia-Echeverria C, Rich DH, Cooper SL. Synthesis of a novel polyurethane copolymer containing covalently attached RGD peptide. J Biomater Sci Poly Ed 1992; 3:217-227. 27. Sugawara T. and Matsuda T. Photochemical surface derivatization of a peptide containing Arg-Gly-Asp (RGD). J Biomed Mater Res 1995; 29:1047-1052. 28. Hubbell JA, Massia SP, Drumheller PD. Surface-grafted cell-binding peptides in tissue engineering of the vascular graft. In: Pedersen H, Matharasan R, DiBiasio D, eds. Annals of the New York Academy of Science 1992:253258. 29. Zdrahala RJ. Small caliber vascular grafts. Part II. Polyurethanes revisited. J Biomater Appl 1996; 11:37-61. 30. Lamba NMK, Woodhouse KA, Cooper SL. Polyurethanes in Biomedical Applications. Boca Raton, FL: CRC Press 1998:277. 31. Annis D, Bornat A, Edwards RO, Higham A, Loveday B, Wilson J. An elastomeric vascular prosthesis. Trans Am Soc Artif Intern Organs 1978; 24:209-214. 32. Gogolewski S, Pennings AJ, Lommen E, Wildevuur CRH, Niewenhuis P. Growth of a neo artery induced by a biodegradable polymeric vascular prosthesis. Makromol Chem Rapid Commun 1983; 4:213-219. 33. Lyman DJ, Alb D, Jackson R, Knutson K. Development of small diameter vascular graft prostheses. Trans Am Soc Artif Intern Organs 1977; 23:253-256. 34. Lyman DJ, Fazzio FJ, Voorhes H, Robinson G, Albo D. Compliance as a factor affecting the patency of a copolyurethane vascular graft. J Biomed Mater Res 1978; 12:337-345. 35. Lin H-B, Sun W, Mosher DF, Garcia-Echeverria C, Schaufelberger K, Lelkes PI, Cooper SL. Synthesis, surface and cell adhesion properties of polyurethanes containing covalently grafted RGD-peptides. J Biomed Mater Res 1994; 28:329-342. 36. Gospardarowicz D, Ill C. Extracellular matrix and control of proliferation of vascular endothelial cells. J Clin Inv 1981; 68:1351-1364. 37. Nichols NK, Gospardarowicz D, Kellser TR, Oslen DB. Increased adherence of vascular endothelial cells to Biomer precoated with extracellular matrix. Trans Am Soc Artif Intern Organs 1981; 27:208-211. 38. Lin H-B, Lewis KB, Leach-Scampavia D, Ratner BD, Cooper SL. Surface properties of RGD-peptide grafted polyurethane block copolymers: Variable take-off angle and cold-stage ESCA studies. J Biomater Sci Polym Ed 1993; 4:183-198. 39. Drumheller PD, Elbert DL, Hubbell JA. Multifunctional poly(ethylene glycol) semiinterpenetrating polymer networks as highly selective adhesive substrates for bioadhesive peptide grafting. Biotech Bioeng 1994; 43:772-780. 40. Massia SP, Hubbell JA. Covalent surface immobilization of Arg-Gly-Asp and Tyr-Ile-Gly-Ser-Arg containing peptides to obtain well-defined cell adhesive substrates. Anal Biochem 1990; 17:292-301.

Matrix Engineering

CHAPTER 51 Hydrogels in Biological Control During Graft Healing Jeffrey A. Hubbell

Introduction—Why Hydrogels?

H

ydrogels are polymeric materials that imbibe a large fraction of water and yet remain intact, not dissolving even given an infinite period of time. These materials are formed from polymer chains that have a high affinity for water, either such that the chains would be individually soluble in water and are restricted from dissolving by virtue of participation in the polymeric material as a network, or such that the chains are almost, but not quite, soluble in water. By network, it is meant that the polymer chains in the hydrogel are somehow interacting with each other so as to keep the individual chains from diffusing away into the aqueous milieu, either by virtue of being covalently bonded together or by interacting physically—specific examples will be provided in the section on “Hydrogel Structure and Synthesis”. Many extracellular structures in the body can be considered as hydrogels. The extracellular matrix of soft tissues and cartilage, for example, exists as a network of glycoproteins and proteoglycans that both interact with each other biophysically (e.g., by specific biological interactions between collagen and a variety of adhesion proteins,1 by specific biological interactions between glycosaminoglycans in proteoglycans and a variety of adhesion proteins,2 and by hydrogen bonding as in collagen due to a very high content of hydroxyproline3) and by covalent interactions (e.g., chemical crosslinking between glutamine residues on one protein and lysine residues on a neighboring protein under the enzymatic action of a transglutaminase to catalyze the formation of an amide linkage between the two).4 It is not by coincidence that hydrogels are found extensively in Nature: They possess distinct biologically useful features, which will likewise be useful in tissue engineering of the vascular graft and other tissues. Hydrogels are mechanically flexible, are freely permeable to small molecules such as dissolved gases and low molecular weight nutrients and wastes, and are controllably permeable to larger molecules such as proteins. Because the gels are highly swollen with water, the amount of polymer mass per unit volume is relatively low, and this is of key importance in cell migration in three dimensions. Cell migration is facilitated by enzymatic remodeling of the three dimensional extracellular matrix, and the amount of protein and proteoglycan mass to be degraded and moved out of the path of the cell is much lower in the case of a hydrogel than if the tissue had consisted of less water. Indeed, in culture models of leukocyte migration through collagen-based gels, cell migration speed was observed to be reduced by nearly an order of magnitude when the collagen concentration was doubled.5 Also, because gels are highly swollen with water, their effective surface area is very high, and this is used in Nature to store large quantities of biologically active molecules that can be released in a triggerable fashion from the extracellular matrix, e.g., the release of heparin-binding polypeptide growth factors from heparan sulfate glycosaminoglycan Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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components of the proteoglycans in the extracellular matrix. As engineered hydrogels are developed for specific applications in tissue engineering, it should be possible to mimic these biological interactions to achieve these same sorts of biological interactions.6-9 Hydrogels have other potential advantages that are less extensively used in Nature. As is addressed in more detail in the section “Protein and Cell Interactions with Hydrogels”, there exists the potential to design gels without the incorporation of binding sites for proteins, with the result being a gel that is highly resistant to cell adhesion. Such materials may be useful in preventing the interaction of cells with material substrates in a scaffold for tissue engineering at sites of application of a gel coating. An additional feature of hydrogels of potential tissue engineering advantage is that a solvent for the gel precursors is water; as such, the possibility exists to apply a hydrogel precursor to a tissue site, dissolved in physiological saline, and to form the hydrogel in situ, as is addressed below in the section “In Situ Transformations”.

Hydrogel Structure and Synthesis Hydrogels consist of hydrophilic polymer chains that are held together in a single mass by virtue of either slight insolubility or some form of bonding between chains to yield a polymer network.10 The structure of the hydrogel is highly dependent upon the nature of the synthetic procedure that was used to obtain the gel; for this reason, different means of hydrogel synthesis will be illustrated in the paragraphs below with a few examples. The nature of some hydrogels is controlled by the slight insolubility of the polymer chains; a typical example is hydrogels of poly(hydroxyethylmethacrylate).11 The vinyl monomer for this polymer, hydroxyethylmethacrylate, is soluble in water. Likewise, oligomers and low molecular weight polymers of hydroxyethylmethacrylate are soluble in water, but high molecular weight polymers are slightly insoluble in water, leading to a gel that swells to imbibe approximately 50% water by mass. This swelling extent can be readily controlled by copolymerization with more hydrophobic monomers, such as ethylmethacrylate (i.e., the analogous monomer to hydroxyethylmethacrylate, only lacking the polar hydroxyl group), to achieve a series of polymers with swelling ranging from almost nil (for poly(ethylmethacrylate)) to that of poly(hydroxyethylmethacrylate). These gels have little internal structure, i.e., no permanent pores or dense regions, but rather exist as dynamic structures with small pores opening and closing in equilibrium. Hydrogels can also be formed by hydrogen bonding, and gels formed from poly(vinyl alcohol) represent a typical example.12 Poly(vinyl alcohol) is a water-soluble polymer at 37°C. When a solution of poly(vinyl alcohol) is dried to form a cast solid, the solid does not dissolve in water when it is again placed in an aqueous milieu. In the dry state, regions of the poly(vinyl alcohol) solid crystallize due to hydrogen bonding. These regions are so strongly bonded that dissolution of the polymer again at 37°C is infinitesimally slow. This behavior can also be obtained, with greater con-

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trol on shape retention, by freezing the poly(vinyl alcohol) solution. When ice crystals form, the polymer chains are excluded from the water crystal and pushed to the crystal boundaries, resulting in very highly concentrated regions of polymer. At these highly concentrated domains, when all of the water is frozen, segments of the poly(vinyl alcohol) chains crystallize due to hydrogen bonding, resulting in an insoluble (without heating far above 37°C) hydrogel in the shape of the polymer solution that was frozen. Thus, this gel is a network polymer where the bonding between polymer chains is physical by hydrogen bonding. Hydrophobic interactions can also be important in the formation of hydrogels, and block copolymers of poly(ethylene glycol) and poly(propylene glycol) form a characteristic example.13 Poly(ethylene glycol) is a watersoluble poly(ether), while poly(propylene glycol) is waterinsoluble, containing a nonpolar methyl group that the poly(ethylene glycol) repeat unit lacks. Block copolymers of these two compositions in the form of a domain of the hydrophobic poly(propylene glycol) flanked on both sides by domains of poly(ethylene glycol) are water soluble polymers that, at relatively high concentrations, can form gels by hydrophobic interactions between the poly(propylene glycol) domains. This is to say, the poly(propylene glycol) domains from several different polymer chains cluster to shield each other from interactions from water, thereby forming physical crosslinks (within these clusters) based on hydrophobic interactions, resulting in a hydrogel. This gel, as do several others based on hydrophobic interactions, has the useful feature that the aqueous mixture is a liquid at cold temperatures (below approximately 10°C) but rapidly forms a hydrogel at 37°C. Electrostatic interactions also play important roles in hydrogel formation, and gels of alginate with calcium ion14 and polyelectrolyte complexes of alginate and poly(lysine)15 form good examples of these principles. Alginate is a naturally occurring polysaccharide product of kelp and several bacterial strains. Each sugar residue in the polysaccharide chain bears a carboxyl group, so the polysaccharide chain is highly anionic, and these charges can play multiple roles in gel formation. Firstly, divalent cations, such as Ca2+, can form a bridge between the anionic groups on adjacent alginate backbones, resulting in regions of adjacent chains that are paired, these regions being connected to other paired regions by amorphous domains. Practically, dropping the alginate solution in Ca2+-free saline into a saline solution containing greater than 1 mM Ca2+ results in the rapid formation of a hydrogel which is stable under physiological conditions. Secondly, alginate (either solution or as a Ca2+crosslinked gel) can be mixed with a polycation in solution, such as poly(lysine), and these two polymer chains form a hydrogel, referred to as a polyelectrolyte complex, crosslinked into a network by the electrostatic interactions between the oppositely charged polymers. Gels formed in either manner described above can have interesting microscopic structures that are lacking in the other gels described in the paragraphs above, this microscopic structure deriving from the diffusive process of gel formation. For example, when the alginate solution is dropped into a Ca2+-containing bath, Ca2+ ions begin diffusing into the polymer solution, causing the

Hydrogels in Biological Control During Graft Healing

formation of the gel at the interface between the polymer solution and the surrounding Ca2+-containing bath. Because the gel region exists at a lower soluble polymer concentration than does the alginate solution, alginate from the core of the droplet diffuses toward the gelled interface; this polymer then gels, and more polymer diffuses from the core to the periphery, and so on: The net result of this phenomenon is that the periphery of the Ca2+-alginate gel is more dense than the core of the particle, owing to the diffusive nature of the process. The case is likewise with the incorporation of the polyelectrolyte counter-ion poly(lysine). Poly(lysine) diffuses into the gel particle from the periphery and interacts first at that periphery, forming denser membrane at the periphery, which restricts further poly(lysine) diffusion into the core in a self-limiting process. Accordingly, a thin shell of polyelectrolyte complex can be formed, a microstructure that could be put to good use in tissue engineering. In addition to the physical interactions described above, hydrogels can also be formed that depend on covalent bonding between polymer chains; these gels have a characteristic advantage over those described above in that they can be considerably more stable under physiological conditions. Three examples are presented below, two based on synthetic polymers with nonbiological reactions and one based on proteins with enzymatic reactions. The simplest way to form a covalently crosslinked hydrogel is to polymerize an otherwise water-soluble polymer in the presence of a multifunctional crosslinker.16 As an example, acrylamide is a monomer that polymerizes to form a water-soluble polymer. If acrylamide is polymerized in the presence of a difunctional monomer (i.e., one that has two sites for polymerization, then referred to as a crosslinker), that monomer will participate in two poly(acrylamide) chains. Accordingly, bis-acrylamido crosslinkers are routinely used to crosslink acrylamide hydrogels. Given that each poly(acrylamide) chain may react with more than one crosslinker molecule, this polymerization process results in the formation of a chemically crosslinked hydrogel. In this case, the entire macroscopic hydrogel object is a single molecule. As a second example of the formation of covalently crosslinked synthetic hydrogels, hydrogels formed from water-soluble polymeric diacrylates form a good example.17 Poly(ethylene glycol) diacrylates (Fig. 51.1)have been formed by coupling one acrylate moiety to each end of a poly(ethylene glycol) chain. These acrylates are capable of polymerizing to form a poly(acrylate), and if a poly(ethylene glycol) monoacrylate were polymerized the resulting polymers would be a comb-like structure, a poly(acrylate) structure forming the back of the comb and poly(ethylene glycol) chains forming the teeth. If, however, a poly(ethylene glycol) diacrylate is polymerized, the acrylates on either end of the chain can enter into different poly(acrylate) chains, resulting in a chemically crosslinked gel that can only be broken apart by the cleavage of a covalent bond. It is also possible for form covalently crosslinked hydrogels using enzymatic crosslinking, and fibrin forms an excellent case in point.18 Fibrin is formed by the enzymatic action of thrombin and factor xiiia on fibrinogen. Throm-

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bin cleaves the fibrinogen protein at two points, exposing a new terminus on two polypeptide chains. This new terminus physically binds at a site on other fibrinogen (and cleaved fibrinogen) proteins, this process resulting in a physically bonded hydrogel. The coagulation protein factor XIII is also cleaved by thrombin to form the active transglutaminase factor xiiia, which acts to catalyze the formation of an amide bond between glutamine residues and lysine residues on opposing polypeptide domains in fibrin, resulting in a chemically crosslinked fibrin network. The brief overview provided above should provide the reader with a background understanding of the physical nature of hydrogels and the vast flexibility in the physical and chemical processes for forming them. Based upon this brief background, the reader is well positioned to consider the application of hydrogels for tissue engineering of the vascular graft.

In Situ Transformations One interesting advantage in the use of hydrogels in tissue engineering is that they may be delivered in the form of a precursor that is dissolved in a physiological saline and is then converted into the insoluble hydrogel upon the target tissue (see Fig. 51.2), in some cases even in direct contact with cells.19 To perform adequately in such an environment, the precursor solution should be lacking in high concentrations of toxic species and low molecular weight monomers, which may be more likely to cross the cell membrane than a high molecular weight macromonomer, and the gel conversion should occur at physiologically acceptable temperatures. This has been accomplished with aqueous solutions of poly(ethylene glycol) diacrylate and related compounds, which may be converted to a hydrogel under biocompatible reaction conditions by exposure to visible light in the presence of a visible light sensitizer.20 The biocompatibility of this process in direct contact with cells can be attributed to the following features: 1. Physiological saline is the solvent; 2. A very high fraction of the polymerizable precursors are macromolecular and thus unable to pass the cell membrane; 3. The photoinitiation process does not involve short wavelength ultraviolet light; 4. The photoinitiation process does not generate large amounts of reactive oxygen species; and 5. The polymerization process does not generate large amounts of heat. These principles serve as a design basis for other biomaterial systems for the formation of hydrogels in situ by transformation from liquid precursors. In situ transformation has at least two major advantages, for some applications, vs. the implantation of preformed hydrogels. One advantage is that a large hydrogel implant may be delivered through a very small surgical hole, somewhat like assembly of the model ship inside the bottle. This has been put toward therapeutic development in the coating of organs in the abdominopelvic cavity following surgery to prevent the formation of postoperative adhesions, engineering the healing response.21 Here is it possible to deliver biomaterial implants as organ coatings

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Fig. 51.1. As an example of designing various mechanisms of degradation and bioactivity of hydrogels, the case of gels formed by crosslinking poly(ethylene glycol) diacrylates is considered. Gels can be made to be substantially non-degradable (a), to incorporate nonenzymatic degradation sites by hydrolysis (b), to incorporate oligopeptide domains (c) which could be sites for cleavage by a protease, e.g. plasmin (d) or which could be sites for binding to adhesion receptors on a targeted cell (e). In each case in this example, the central block of the macromer is a poly(ethylene glycol) chain, selected for resistance to protein adsorption and for favorable toxicology of the degradation product (f). These diacrylated poly(ethylene glycol) chains polymerize by free radical reaction at the acrylate to form oligomers of acrylic acid that are linked to other oligomers of acrylic acid via the poly(ethylene glycol) chains with the relevant degradation sites (g). Thus, the acrylic acid oligomers form nodes in a three dimensional network, each connected to other nodes by long strings of poly(ethylene glycol). The permeability of the gels to proteins, and the rate of release of incorporated protein drugs, can be controlled via manipulation of the density of the oligomeric acrylic acid nodes, by the number of poly(ethylene glycol) strings emanating therefrom, and by the length of the poly(ethylene glycol) strings. Degradation products are poly(ethylene glycol) and oligomers of acrylic acid, which are eliminated by glomerular filtration, and lactic acid.

that are many centimeters in diameter, but to deliver these implants using laparoscopic instruments to convert the liquid precursor to a hydrogel. A second advantage is that hydrogel implants of very complex shape may be formed in situ; this has been put to advantage in the coating of arterial surfaces after deendothelialization to prevent the deposition of platelets on the subendothelium.22,23 In this case, the photoinitiator was separated from the other components of the precursor solution and was adsorbed to the arterial surface; when the arterial lumen was flooded with all of the remaining precursors and the artery was irradiated with visible light, the formation of the hydrogel was restricted to the interface between the arterial wall and its contents. Thus, a hydrogel barrier was formed, only several microns thick, that was conformal with the complex shape of the deendothelialized artery. Given that one may want to employ hydrogels in tissue engineering for the delivery of therapeutics and even living cells to polymeric scaffolds and even preexisting ar-

teries, to attempt to engineer a healing response, the ability to form gels in situ presents an attractive feature. Such transformations can be accomplished even in the presence of individual and aggregated cells,24 suggesting that such chemical schemes can be adapted for the delivery of cellular therapeutics either within or adjacent to the vascular system.

Protein and Cell Interactions with Hydrogels Hydrogels can be designed to have minimal interactions with proteins and cells. Cell adhesion to biomaterials is mediated by an adsorbed protein overlayer.25 Protein adsorption to materials is dominated primarily by hydrophobic interactions and secondarily by electrostatic interactions. 26 Most insoluble polymers (e.g., poly(ester)s, poly(ethylene), poly(propylene), poly(urethane)s, and poly(tetrafluoroethylene)) achieve their water-insolubility by interacting poorly with water. These hydrophobic polymers accordingly adsorb large amounts of proteins, which act as surfactants to expose more hydrophobic domains to-

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Fig. 51.2. As an example of a transformation that can be carried out in situ, hydrogels can be formed from macromolecular precursors dissolved in buffered saline, directly on the tissue surface. A tissue surface to be treated (a), such as the anastomotic region of a vascular graft, is first flushed with a solution of a nontoxic photoinitiator (in this example eosin Y), to stain the tissue surface with the photoinitiator (b). After the nonadsorbed photoinitiator is flushed away and the artery lumen is filled with the other components of the initiation system (in this example, the nontoxic buffer triethanolamine) and the polymerizable macromer (such as a diacrylated poly(ethylene glycol) copolymer with lactic acid, as in Fig. 51.1) (c), the artery is irradiated with light at an appropriate wavelength (in the case of eosin Y, green light) to cause a wave of polymerization to proceed from the tissue surface out into the lumen (d). After the nonpolymerized macromer solution is rinsed away, the final result is a crosslinked solid hydrogel, adherent to and conformal with the tissue surface (e). In the photoinitiation scheme shown in this example, the red eosin Y adsorbs green light and is excited to a triplet state, from which it participates in a single-electron transfer reaction with triethanolamine, resulting in a radical on the eosin and the triethanolamine. The free radical on the triethanolamine rearranges and serves as the actual initiator of the free radical polymerization of the diacrylated macromer.

ward the material surface and thus present a more hydrophilic new surface to the aqueous surroundings. Electrostatic interactions are secondary (due to the high ionic strength of physiological fluids) but remain important in protein interaction with materials, especially with regard to adsorption of proteins to cationic polymers (since almost all proteins bear a net anionic charge). Accordingly, one may design a biomaterial to be relatively free of cell adhesion by making it be both very hydrophilic (e.g., a hydrogel) and nonionic or anionic. Much of what is understood about protein and cell interactions with hydrogels has been learned from two dimensional hydrogel models, namely water-soluble polymers (which would form hydrogels) tethered to insoluble surfaces.

Work on the theory of these interactions has been performed by Jeon et al.27,28 These investigators determined that when the tethered polymer brush exists at a sufficient density of polymer chains of a sufficient length, compression of the polymer brush by a potentially adsorbing protein results in a strong resistance, i.e., steric stabilization, to the adsorption of the protein. Thus, one may understand with these theoretical analyses based on the physics of polymer chains that the thermodynamic interactions are unfavorable for protein adsorption.29 Very well characterized experimentation has also been performed with model two dimensional polymer brushes, e.g., by Prime and Whitesides employing self-assembling monolayers of alkane thiols on gold substrates, the alkane thiols bearing grafted poly(ethylene glycol)

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chains.30 These studies demonstrated experimentally that a high density of poly(ethylene glycol) chains would suffice, even if the chains were very short, to dramatically limit protein adsorption. The physical character of hydrogels is important, in tandem with the chemical character, in determining hydrogel biocompatibility. This is to say, it is not just the chemical structure of the repeat unit of the polymer in the hydrogel that determines the biocompatibility of the hydrogel. This can be considered in the context of crosslinked gels of poly(ethylene glycol) diacrylate.31 When gels were formed from poly(ethylene glycol) of molecular weight 200 or 1000, these materials were rapidly encapsulated in a typical foreign body reaction and calcified after subcutaneous implantation in the rat; however, gels formed from poly(ethylene glycol) diacrylate of higher molecular weight remained free of fibrous encapsulation and did not calcify. The calcification response may be due to a specific interaction between Ca2+ and poly(ethylene glycol), but the induction of a foreign body reaction almost certainly relates to the difference in water content of these gels (from approximately 75% water for the poly(ethylene glycol) 200 diacrylate to approx. 90% for the higher molecular weight poly(ethylene glycol)-based hydrogels. As such, one should understand clearly that all hydrogels are not alike, and even that all hydrogels of the same chemical composition are not necessarily alike in their biological responses.

Designed Cell Adhesiveness of Hydrogels In the above section it was demonstrated that, with adequate attention to the details of hydrogel design, hydrogels can be formed to effectively resist the attachment of cells, even in vivo. Based upon this observation it then becomes attractive to incorporate into such a hydrogel biological adhesion ligands targeting specific receptors and even perhaps specific cell types. In this manner, one might be able to develop a hydrogel material for tissue engineering that would reject the attachment on one cell type (e.g., the platelet) but that would accept the attachment of another cell type (e.g., the endothelial cell). Tissue engineering with synthetic adhesion ligands is based on profound advances in the molecular biology of cell adhesion over the past several years, in which a host of extracellular matrix adhesion glycoproteins have been identified, together with their corresponding receptors. Moreover, the domains on many of the adhesion proteins that bind to the receptors have been identified, and it has been demonstrated that, in many of these cases, synthetic peptide and nonpeptide analogs can be developed with appropriate affinity for the appropriate receptors on the cell surfaces. It has also been demonstrated that some of these adhesion ligands have affinity for receptors that exist on one cell type do not exist on other cell types present in the same biological environment; e.g., the sequence REDV, which is present in some fibronectin forms, binds to the integrin receptor α4β1 that is present on human endothelial cells but not on human blood platelets.32,33 The biology of adhesion proteins, their receptors and synthetic analogs of the adhesion proteins is the topic of chapter 54 in this book and is not addressed further here; only the

Tissue Engineering of Prosthetic Vascular Grafts

biomaterials engineering aspects of incorporation of these signals into hydrogels are addressed. In order to be able to engineer a selective adhesion response to hydrogel materials, one must devise means by which to incorporate synthetic adhesion ligands, such as synthetic peptides derived from adhesion proteins, into hydrogel networks. Hydrogels derived from poly(ethylene glycol) diacrylate again serve as an instructive example. Hern and the author34 examined means by which to incorporate adhesion peptides into these photopolymerizable gels, using the adhesion peptide GRGDY as a model (the N-terminal G serving as a spacer, and the C-terminal Y serving as a site for radioiodination). The N-terminus of the peptide was acrylated, either with a minimal spacer or with a poly(ethylene glycol) chain of molecular weight 3400, and these acrylated peptides were incorporated into the hydrogels by copolymerization (i.e., mixing the peptide acrylates and the poly(ethylene glycol) diacrylates in the hydrogel precursor solution, and then photopolymerizing the mixture). When the peptides were omitted from the hydrogel, the resulting material was very resistant to cell adhesion. When an acrylated RGD or control inactive RDG peptide was incorporated without a spacer arm, the resulting material supported cell adhesion in both cases, but only in the presence of serum proteins; this result can only be explained by the peptide being sterically unavailable for biospecific binding to cell adhesion receptors but serving as a site for the nonspecific adsorption of serum proteins. When the RGD and RDG peptides were incorporated with a poly(ethylene glycol) spacer, the materials with RGD supported cell adhesion, while those with the inactive RDG supported no adhesion. Thus, it would seem that for the sequence to be sterically available for binding to receptors on the cell it must be immobilized within the gel via a long spacer; and it would further seem that the peptide immobilized without the spacer serves as a site for protein adsorption, while the peptide immobilized with the spacer does not. Having these results in hand, one is then prepared to begin to incorporate cell type specific adhesion signals (such as the endothelial targeting REDV sequence) to engineer the tissue response to a hydrogel, and thus to a vascular graft.

Nonenzymatic Degradation of Hydrogels It is clear that one must have in one’s repertoire of materials for tissue engineering hydrogels that are degradable. Degradation by nonenzymatic hydrolysis can be obtain by incorporating hydrolytically labile bonds within the gel precursor, either in a crosslinker or within the polymer backbone. The example of poly(ethylene glycol) diacrylate hydrogels, in this case degradable variations, will be selected for discussion. A host of degradable polymers are available for use in medicine, and these are based on hydrolysis of esters, carbonates, and anhydrides, to name only a few of the important approaches (see chapter 47). These compositions can be incorporated into polymers that form hydrogels, to obtain the physical and biological properties of the hydrogel along with the degradability inferred by the water-labile bond. To accomplish this in the scheme of photopolymerized poly(ethylene glycol) diacrylate hydrogels, block copolymers

Hydrogels in Biological Control During Graft Healing

have been formed, with poly(ethylene glycol) forming the central block and oligomers of, e.g., lactic acid forming flanking blocks; the termini of the lactic acid blocks can be acrylated, leading to an analog of poly(ethylene glycol) diacrylate with a degradable segment between the poly(ethylene glycol) segments.35 The hydrogel to result from such a precursor consists of nodes of oligo(acrylic acid) esterified to poly(ethylene glycol) chains via an oligo(lactic acid) intermediate. Degradation thus yields a low molecular weight oligo(acrylic acid), lactic acid and its oligomers, and poly(ethylene glycol) chains of the original molecular weight. The rates of degradation of hydrogels behave somewhat differently than those of the poly(ester), poly(carbonate), or poly(anhydride) parent family. In the parent families, two features control the rate of degradation: The intrinsic susceptibility of the water-labile bond to hydrolysis, and the rate of transport of water into the hydrophobic polymer. In hydrogels, by contrast, the entire macroscopic material is equally exposed to water, so the intrinsic rate of hydrolysis controls the overall degradation rate. This feature can be used to modulate the rate of resorption of such hydrogels: For example, poly(ethylene glycol)glycolide diacrylate hydrogels degrade faster than poly(ethylene glycol)-lactide diacrylate hydrogels, which degrade faster than poly(ethylene glycol)-caprolactone diacrylate hydrogels, in accordance with the expected order based on chemical kinetics.

Enzymatic Degradation of Hydrogels In addition to degradation based on nonenzymatic hydrolysis, one could contemplate degradation based on enzymatic processes. Enzymatic degradation may have conceptual advantages over nonenzymatic degradation in some applications of tissue engineering. For example, cell migration through tissues and through fibrin clots and granulation tissue in wound healing is dependent upon enzymatic hydrolysis of glycoproteins and glycosaminoglycans in the extracellular matrix. Thus, endowing a hydrogel with the ability to be degraded by enzymes would permit the hydrogel to be remodeled by biological processes associated with cell migration. It is instructive to consider cell migration through biologically derived hydrogels, as a prototype for the ultimate design of synthetic materials that permit cell infiltration by proteolysis. Fibrin serves as one such good prototype, being the primary matrix through which cells migrate in a healing response. Herbert et al examined the migration of the growth cones of neurites emanating from aggregated neural and glial cells within a three dimensional fibrin matrix in a culture model.36,37 Fibrin was formed under controlled conditions around the cell aggregates, and the rate of neurite extension was measured as a function of fibrin fibril density and morphology; additionally, the dependence of the extension process on local protease activity was examined. It was observed that either direct inhibition of plasmin activity or inhibition of plasminogen activation activity completely inhibited cell migration, indicating that the neurites did not find preexisting pores through which to migrate, and that they did not mechanically push fibrin fibrils out of their path during migration, but rather that they expressed chemical

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activity to clear the fibrin from their path. Electron microscopic observation of the morphology of the neurite tips and the fibrin near the neurite tips demonstrated that fibrils were present in the immediate vicinity of all aspects of the growth cone, with less than a 50 nm gap between the neurite plasma membrane and the fibrin fibrils, suggesting that the neurite did not migrate by finding a pathway that was enzymatically created in a spatially nonlocalized manner, but rather that local enzymatic activity associated with the cell membrane was responsible for local removal of material to create a pathway. Furthermore, these studies also demonstrated that there was no dependence in neurite migration rate on fibrin fibril morphology, but only on density; i.e., the only feature that mattered was the amount of material present to be cleared from the pathway of the cell to permit migration. Studies as described above both emphasize the importance of biologically derived materials for tissue engineering and also provide motivation to develop synthetic analogs. The fact that, at least in some circumstances, the detailed morphology of the biological structures to be cleared by cell-associated proteolytic activity is less important than the mass of that material encourages one to proceed with development of such synthetic analogs, in that only the chemical nature need be mimicked, not the morphological nature as well (at least in some cases). With this in mind, West et al have examined poly(ethylene glycol)-based hydrogel materials that degrade by proteolytic activity.38 Poly(ethylene glycol) was used as a substrate for liquid phase peptide synthesis, and either plasmin or collagenase-sensitive peptide sequences were synthesized at the termini of the polymer chains. Upon the termini of these peptide blocks an acrylate group was coupled, to create an analog of the poly(ethylene glycol)-lactide diacrylate, namely poly(ethylene glycol)-peptide diacrylate, where the identify of the peptide was selected for sensitivity to the protease plasmin or collagenase. Hydrogels were formed from such precursors and were exposed to either plasmin or collagenase, with the result that the plasmin-sensitive gel degraded in the presence of plasmin but not collagenase, and the collagenase-sensitive gel degraded in the presence of collagenase but not plasmin, as designed. Such preliminary studies demonstrate at least the potential for the synthesis of hydrogel materials that mimic the ability to be remodeled like the natural biological hydrogels used in nature.

Controlled Release Hydrogels have been extensively explored for the controlled release of drugs, particularly macromolecular drugs such as proteins and oligonucleotides. One reason for utilizing hydrogels for drug delivery is the relative ease in controlling the permeability of the gel structure to the macromolecular drug. While in most gels permanent pore structures do not exist, transient pores do exist that can be described in terms of molecular weight between crosslinks and pore dimensions.39 These parameters can be readily controlled, and by adjusting these dimensions relative to the radius of gyration of the macromolecular drug one can either restrict drug diffusion to a controlled degree and obtain release from the hydrogel material on a time scale faster

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than that of degradation, or one can entrap the drug, using a gel with a pore dimension much less than the drug size, and obtain drug release that is mediated by the rate of hydrogel degradation.40 It is relatively straightforward to synthetically manipulate the pore dimensions of a hydrogel. For example, with gels formed by crosslinking an unsaturated monomer (refer to the example of acrylamide with a bisacrylamido crosslinker) the pore dimensions may be altered via control of the ratio of acrylamide to bis-acrylamide, a high amount of bis-acrylamido crosslinker yielding small pore dimensions and restricted permeability. In the example of poly(ethylene glycol)-diacrylate based hydrogels, one may alter pore dimensions via the molecular weight of the poly(ethylene glycol) diacrylate, low molecular weights yielding small pore dimensions and low permeability. As an example, Cruise et al41 reported pore dimensions of 19 Å for gels formed from poly(ethylene glycol) 2000 diacrylate, 22 Å from poly(ethylene glycol) 4000 diacrylate, 34 Å from poly(ethylene glycol) 8000 diacrylate and 58 Å from poly(ethylene glycol) 20,000 diacrylate. It was likewise possible to control pore dimensions via manipulation of the reaction conditions with the same poly(ethylene glycol) diacrylate precursor; e.g., pore dimensions of 70 Å were obtained by polymerization of poly(ethylene glycol) diacrylate 20,000 from a 10% precursor solution, 58 Å from a 20% precursor solution, and 42 Å from a 30% precursor solution. Because of this high flexibility in the control of pore dimensions and gel permeability to macromolecular drugs, hydrogels are extremely attractive for controlled release applications in tissue engineering.

Design Principles, Hydrogels for Biological Control During Graft Healing One may consider many possible applications for hydrogels in tissue engineering, some of which have been explored already and some of which remain to be experimentally addressed. In this section, some of the more important possible applications in vascular tissue engineering will be called into focus. These relate to altering the blood response to a graft, altering the arterial response to the graft, and directing the response of the surrounding tissue to a graft implant. Reduction of Thrombosis Because hydrogels can be designed that adsorb relatively low amounts of protein and that thus support relatively low amounts of platelet adhesion (which has been demonstrated to be modulated principally by adsorbed fibrinogen),42 one obvious, yet important, application of hydrogels to vascular graft tissue engineering is the reduction of platelet deposition in the hours and days following implantation. Hydrogel coatings have been used to modulate the deposition of platelets and other cells on a host of medical devices, most of which are implanted for relatively short periods of time. Indeed, heparinization of graft surfaces may improve graft biocompatibility by both a pharmacological effect (the specific biological activity of heparin as it interacts with antithrombin III) as well as a biophysical effect (the hydrophilic, anionic heparin hydrogel coating acting to reduce the adsorption of fibrinogen).43

Tissue Engineering of Prosthetic Vascular Grafts

Graft Sealing Knitted and woven grafts must be preclotted to prevent excessive blood leakage from the graft, and this process of preclotting may predispose the graft to further thrombosis in the short term (e.g., due to platelet interactions with fibrin in the clot or due to prothrombin activation on the surfaces of cells in the clot) and to the development of an excessive neointima in the long term (e.g., due to the growth factors that are released by activated platelets in the clot). To reduce the possibility of short and long term graft failure as a result of preclotting, grafts that are based on collagen,44 gelatin,45 and albumin46 preclotting have been developed. These materials are indeed hydrogels, and thus hydrogelsealed vascular grafts have already been used clinically. One could also employ hydrogels designed more specifically for graft sealing. Such a hydrogel sealing material could either resorb by physical erosion (e.g., with poly(ethylene glycol)poly(propylene glycol)-poly(ethylene glycol) block copolymers) or by nonenzymatic hydrolysis (e.g., with a poly(ethylene glycol)-lactide diacrylate hydrogel). Such a designed hydrogel has the possible benefit that other features could also be designed into the gel, as described below. Promotion of Endothelialization It is clear that the rapid establishment of an endothelial cell monolayer atop the neointima within a graft is key to success in the tissue engineering of a vascular graft. It is also clear that thrombosis within the graft lumen is an undesirable outcome. In light of these two observations, one could consider employing a hydrogel coating or a hydrogel sealing to both reduce thrombosis and to enhance endothelialization, e.g., via a gel that resists fibrinogen adsorption and accordant platelet deposition, that gel also containing a grafted ligand that demonstrates specificity for an endothelial cell adhesion receptor. In such a manner, it might be possible to fail to induce platelet adhesion but to succeed in inducing endothelial cell adhesion.32 In the absence of seeding of endothelial cells along the entire lumen of the graft, one must be concerned with the rate of endothelial cell migration along the coated graft surface. Lauffenburger and colleagues have demonstrated, both based on theoretical considerations and also by experiment, that there exists an optimum in the surface density of adhesion ligands for fastest cell migration rates.47 At low surface density of adhesion ligand, cell traction is low and cell migration rates, as well as adhesion strength, are accordingly slow. At high surface density of adhesion ligand, cell traction is very high, but the rate of binding of an adhesion receptor on the cell surface with surface-bound adhesion ligand is very high, resulting in very strong adhesion but very low cell migration rates. Accordingly, one must design the display of adhesion ligands in the hydrogel coating such that adhesion strength is sufficient to resist fluid forces within the vascular flow but such that it is not so strong that the rate of cell migration is unacceptably reduced. It has been demonstrated that thrombosis at the anastomoses of a graft may be well in excess of thrombosis on the graft itself.48 Accordingly, efforts at reducing platelet deposition on the biomaterial per se may be slightly mis-

Hydrogels in Biological Control During Graft Healing

placed relative to the thrombogenic potential of the anastomoses. One may then consider the use of a hydrogel to passivate the graft anastomoses. It was demonstrated above that one advantage of hydrogels over other materials for tissue engineering is that they may be formed on the surface of a tissue by an in situ transformation. As an example, it was discussed that hydrogel coatings have been formed on the surfaces of injured arteries to replace one function of the endothelium, namely to regulate the deposition of blood platelet on the arterial surface.22 Accordingly, it could be envisioned to limit the extent of thrombus deposition on the anastomoses within a graft that had already been sutured in place by the local, in situ deposition of a hydrogel coating spanning the junction of the native artery and the implanted graft. Using such an approach, it might be possible to directly impact one of the most troublesome regions of an implanted graft. Drug Release Hydrogels have been explored extensively in drug release, due to the facility with which the permeability of the gel material to a drug of a given size may be modulated. Because of this high flexibility in control of release characteristics, it is attractive to combine some of the other uses described above with local release of a biologically active factor, e.g., an angiogenic growth factor such as basic fibroblast growth factor or vascular endothelial growth factor, to induce a specific biological response in tissue engineering of the vascular graft. Some experimentation has been performed along these lines, using hydrogels in vascular tissue engineering. For example, Simons et al49 employed a hydrogel, formed in situ by application of liquid solution of a copolymer of poly(ethylene glycol)-poly(propylene glycol)poly(ethylene glycol) block copolymer, which solidifies to form a gel upon warming to body temperature.13 This gel was used to release an antisense oligonucleotide that inhibited cell proliferation; this treatment was observed to greatly limit the thickening of the neointima following injury by balloon angioplasty of the carotid arteries of rats. As another example, Greisler and colleagues50 employed a fibrin-based hydrogel formed within the pore spaces of a vascular graft to release basic fibroblast growth factor, and this resulted in an overall acceleration of the rate of endothelialization and a corresponding limitation of the formation of an excessively thickened neointima in animal models. Promotion of Transmural Capillary Ingrowth Studies by Golden et al provided great encouragement for the possibility of transmural capillary infiltration and differentiation to form large vessel endothelium with large pore graft implants in nonhuman primate models.51 One can consider designing hydrogel treatments for tissue-engineered vascular grafts to attempt to achieve similar results in the human. In such an approach, one might integrate the concepts that have been presented above: For example, one might consider sealing a graft with a hydrogel that bore cell adhesion sites for endothelial cells but not for blood platelets, which further bore sites for degradation by proteases expressed during endothelial cell migration. One might further incorporate a growth factor within such a hydrogel to

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stimulate endothelial cell migration and proliferation, and another factor to limit perigraft inflammation or smooth muscle activity. Stated more generally, Nature’s extracellular matrix hydrogel serves to provide a great deal of information to cells to direct their behavior and organization in space. It provides information in the form of structural cues, of immobilized adhesion molecules, and of immobilized and diffusible polypeptide growth factors. Furthermore, it receives information from the cells as they respond. It does this by binding to growth factors that are released from cells and by degrading in the face of cellularly-displayed enzymes. It might be possible, and indeed advantageous, in tissue engineering of the vascular graft, to mimic some or all of these activities, either in a biologically derived hydrogel matrix or in a completely synthetic biomimetic hydrogel matrix, with the goal of harnessing some of these biological activities for controlling tissue morphogenesis and function. References 1. Shimizu M, Minakuchi K, Moon M, and Koga J, Difference in interaction of fibronectin with type I collagen and type IV collagen. Biochim Biophys Acta 1997; 1339:53-61. 2. Jackson RL, Busch SJ, and Cardin AD, Glycosaminoglycans: Molecular properties, protein interactions, and role in physiological processes. Physiol Rev 1991; 71:481-539. 3. Vanderrest M, and Garrone R, Collagen family of proteins. FASEB J 1991; 5:2814-2823. 4. Mosher DF, Assembly of fibronectin into extracellular matrix. Curr Opin Struct Biol 1993; 3:214-222. 5. Parkhurst MR, and Saltzman WM, Quantification of human neutrophil motility in 3-dimensional collagen gels: Effect of collagen concentration. Biophys J 1992; 61:306-315. 6. Langer R, and Vacanti JP, Tissue engineering. Science 1993; 260:920-926. 7. Peppas NA, and Langer R, Challenges in biomaterials. Science 1994; 263:1715-1720. 8. Hubbell JA, Biomaterials in tissue engineering. Bio/Technology 1995; 13: 565-576. 9. Elbert DL, and Hubbell JA, Surface treatments of polymers for biocompatibility. Ann Rev Material Sci 1996; 26: 365-394. 10. Kim SW, Bae YH, and Okano T, Hydrogels: Swelling, drug loading, and release. Pharm Res 1992; 9:283-290. 11. Slack SM, and Horbett TA, Changes in fibrinogen adsorbed to segmented polyurethanes and hydroxyethylmethacrylate-ethylmethacrylate copolymers. J Biomed Mater Res 1992; 26:1633-1649. 12. Mallapragada SK, Peppas NA, and Colombo P, Crystal dissolution-controlled release systems. II. Metronidazole release from semicrystalline poly(vinylalcohol) systems. J Biomed Mater Res 1997; 36:125-130. 13. Leach RE, and Henry RL. Reduction of postoperative adhesion in the rat uterine horn model with Poloxamer 407. Am J Obstet Gynecol 1990; 162:1317-1319. 14. O’Shea GM, and Sun AM, Encapsulation of rat islets of Langerhans prolongs xenograft survival in diabetic mice. Diabetes 1986; 35:943-946. 15. Decher G, Lvov Y, and Schmitt J, Proof of multilayer structural organization in self-assembled polycationpolyanion molecular films. Thin Solid Films 1994; 244:772-777.

570 16. Gin H, Dupuy B, Caix J, Baquey C, and Ducassou D, In vitro diffusion in polyacrylamide embedded agarose microbeads. J Microencapsul 1990; 7:17-23. 17. Pathak CP, Sawhney AS, and Hubbell JA, Rapid photopolymerization of gels in contact with cells and tissue. J Am Chem Soc 1992; 114:8311-8312. 18. Moser D, Blood coagulation and fibrinolysis. Clin Cardiol 1990; 13:5-11. 19. Hubbell JA, In situ material transformations in tissue engineering. MRS Bulletin 1996; 21:33-35. 20. Sawhney AS, Pathak CP, Van Rensberg JL, and Hubbell JA, Optimization of photopolymerized bioerodible hydrogel properties for adhesion prevention. J Biomed Mater Res 1994; 28: 831-838. 21. Hill-West JL, Chowdhury SM, Sawhney AS, Pathak CP, Dunn RC, and Hubbell JA, Prevetion of postoperative adhesions in the rat by in situ photopolymerization of bioresorbable hydrogel barrires. Obstet Gynecol 1994; 83:59-64. 22. Hill-West JL, Chowdhury SM, Slepian MJ, and Hubbell JA, Inhibition of thrombosis and intimal thickening by in situ photopolymerization of thin hydrogel barriers. Proc Nat Acad Sci USA 1994; 91:5967-5971. 23. West JL and Hubbell JA, Separation of the arterial wall from blood contact using hydrogel barriers reduces intimal thickening after balloon injury in the rat: The roles of medial and luminal factors in arterial healing. Proc Nat Acad Sci USA, 1996; 93:13188-13193. 24. Sawhney AS, Pathak CP, and Hubbell JA, Modification of islet of Langerhans surfaces with immunoprotective poly(ethylene glycol) coatings via interfacial photopolymerization. Biotechnol Bioeng 1994; 44:383-386. 25. Pettit DK, Hoffman AS, and Horbett TA, Correlation between corneal epithelial cell outgrowth and monoclonal antibody binding to the cell-binding domain of adsorbed fibronectin. J Biomed Mater Res 1994; 28:685-691. 26. Andrade JD, and Hlady V, Protein adsorption and materials biocompatibility: A tutorial review and suggested hypotheses. Adv Polym Sci 1986; 79:1-63. 27. Jeon Si, Lee JH, Andrade JD, and de Gennes PG, Protein-surface interactions in the presence of polyethylene oxide I, simplified theory. J Coll Interf Sci 1991; 142:149-158. 28. Jeon SI, and Andrade JD, Protein-surface interactions in the presence of polyethylene oxide II, effect of protein size. J Colloid Interface Sci 1991; 142:159-166. 29. Norde W, Protein adsorption at solid surfaces: A thermodynamic approach. Pure Appl Chem 1994; 66:491-496. 30. Prime KL, and Whitesides GM, Adsorption of proteins onto surfaces containing end-attached oligo(ethylene oxide): A model system using self-assembled monolayers. J Am Chem Soc 1993; 115:10714-10721. 31. Hossainy SFA, and Hubbell JA, Molecular weight dependence of calcification of poly(ethylene glycol) hydrogels. Biomaterials 1994; 15:921-925. 32. Hubbell JA, Massia SP, Desai NP, and Drumheller PD Endothelial cell-selective materials for tissue engineering in the vascular graft via a new receptor. Bio/Technology 1991; 9:568-572. 33. Massia SP, and Hubbell JA, Vascular endothelial cell adhesion and spreading promoted by the peptide REDV of the IIICS region of plasma fibronectin is mediated by integrin α4β1. J Biol Chem 1992; 267:14019-14026.

Tissue Engineering of Prosthetic Vascular Grafts 34. Hern DL, and Hubbell JA, Incorporation of adhesion peptides into nonadhesive hydrogels useful for tissue resurfacing. J Biomed Mater Res, 1997; 39:266-276. 35. Sawhney AS, Pathak CP, and Hubbell JA, Bioerodible hydrogels based on photopolymerized poly(ethylene glycol)-co-poly(α-hydroxy acid) diacrylate macromers. Macromolecules 1993; 26:581-587. 36. Herbert CB, Bittner GD, and Hubbell JA, Effects of fibrinolysis on neurite growth from dorsal root ganglia cultured in 2-dimensional and 3-dimensional fibrin gels. J Comp Neurol 1996; 365:380-391. 37. Herbert CB, Nagaswami C, Bittner GD, JW Weisel, and Hubbell JA, Effects of fibrin micro-morphology on neurite growth from dorsal root ganglia cultured within threedimensional fibrin gels. J Biomed Mater Res, 1997; in ress. 38. West JL, Pratt A, and Hubbell JA, Proteolytically degradable hydrogels, Trans Soc Biomaterials 1997; 23:103. 39. Canal T, and Peppas NA, Correlation between mesh size and equilibrium degree of swelling of polymeric networks. J Biomed Mater Res 1989; 23:1183-1193. 40. JA Hubbell, Hydrogel systems for barriers and local drug delivery in the control of wound healing. J Contr Rel 1996; 39:305-313. 41. Cruise GM, Scharp DS, and Hubbell JA, Characterization of permeability and network structure of interfacially photopolymerized poly(ethylene glycol) diacrylate hydrogels. Biomaterials 1998; 19:1287-1294. 42. Kiaei D, Hoffman AS, Horbett TA, and Lew KR, Platelet and monoclonal antibody binding to fibrinogen adsorbed on glow-discharge-deposited polymers. J Biomed Mater Res 1995; 29:729-739. 43. Hubbell JA, Pharmacologic modification of materials. Cardiovasc Pathol 1993; 2:121S-128S. 44. Tolan M, Wells F, Kendall S, Large S, and Wallwork J, Clinical experience with a collagen-impregnated woven Dacron graft. J Cardiovasc Surg 1995; 36:323-327. 45. Utoh J, Goto H, Obayashi H, Hirata T, and Miyauchi Y, Dilation of gelatin-sealed knitted dacron prosthesis. J Cardiovasc Surg 1996; 37:343-344. 46. Kang SS, Petsikas D, Murchan P, Cziperle DJ, Ren D, Kim DU, and Greisler HP, Effects of albumin coating of knitted Dacron grafts on transinterstitial blood loss and tissue ingrowth and incorporation. Cardiovasc Surg 1997; 5:184-189. 47. DiMilla PA, Stone JA, Quinn JA, Albelda SM, and Lauffenberger DA, Maximal migration of human smooth muscle cells on fibronectin and type IV collagen occurs at an intermediate attachment strength. J Cell Biol 1993; 122:729-737. 48. Johnson PC, Dickson CS, Garrett KO, Sheppeck RA, and Bentz ML, The effect of microvascular anastomosis configuration on initial platelet deposition. Plast Reconstr Surg 1993; 91:522-527. 49. Simons M, Edelman ER, DeKeyser JL, Langer R, and Rosenberg RD, Antisense c-myb oligonucleotides inhibit intimal arterial smooth muscle cell accumulation in vivo. Nature 1992; 359:67-70. 50. Greisler HP, Cziperle DJ, Kim DU, Garfield JD, Petsikas D, Murchan PM, Applegren EO, Drohan W, and Burgess WH, Enhanced endothelialization of expanded polytetrafluoroethylene grafts by fibroblast growth factor type 1 pretreatment. Surgery 1992; 112:244-254. 51. Golden MA, Hanson SR, Kirkman TR, and Clowes AW, Healing of poly(tetrafluoroethylene) arterial grafts is influenced by graft porosity. J Vasc Surg 1990; 22:838-844.

Matrix Engineering

CHAPTER 52 In Vivo Synthesis of Organs Using Collagen-GAG Copolymers Ioannis V. Yannas

Introduction

A

s practiced today, the methodology of tissue engineering emphasizes the reconstruction of tissues and organs. Such reconstructive activity may be either induced to take place inside the host organism, at the site of the missing organ (in vivo synthesis), or else it may be induced outside the organism (in vitro synthesis), typically in a cell culture environment, prior to being grafted onto the host. In vivo synthesis of organs has been induced by use of analogs of the extracellular matrix (ECM). Three organs have been synthesized so far by this method, namely, the skin (in humans, as well as in swine and guinea pig models), peripheral nerves (rat) and the knee meniscus (canine). The procedure depends critically on identification of the appropriate regeneration template, a biologically active ECM analog which can block scar synthesis and induce instead the desired in situ synthesis of the physiological organ. Convincing evidence that regeneration has, in fact, occurred must be based on observation of definitive recovery of both structure and function in an anatomically well defined defect. Furthermore, there must be evidence that the mass of the regenerated organ which forms in the absence of the template (spontaneous regeneration) is negligibly small relative to that which is induced in its presence. In this chapter, I will first present the evidence supporting the conclusion that organ regeneration can indeed be reproducibly achieved both in animal models and clinically. The methodology of induced regeneration will then be briefly summarized, followed by an introduction to the models which are used to describe the mechanism of induced regeneration. Although in vivo synthesis of vascular tissue has not yet been demonstrated, I suggest that future experimentation in this area could be very fruitful and could benefit directly by use of the principles described briefly below for other tissues and organs.

Summary of Evidence for In Vivo Synthesis of Organs There is considerable evidence that the adult mammalian dermis does not regenerate spontaneously.1,2 In studies which have been conducted since 1970 it has have shown that a porous graft copolymer of type I collagen and chondroitin 6-sulfate, a glycosaminoglycan (collagen-GAG copolymer) induces in vivo synthesis (induced regeneration) of the dermis Tissue Engineering of Prosthetic Vascular Grafts, edited by Peter Zilla and Howard P. Greisler. ©1999 R.G. Landes Company.

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in large areas of full-thickness skin loss in the guinea pig.3-10 This finding has been independently confirmed repeatedly in humans11-13 and recently in swine.14 The evidence for regeneration was based on histological and ultrastructural studies, the use of small-angle laser light scattering studies of histological sections, and functional studies. It was concluded that the new skin was structurally and functionally competent; however, it was totally lacking in hair follicles and other skin adenexa. The emphasis in early studies of in vivo synthesis of skin was on keratinocyte seeding of the analog prior to grafting, as this procedure led to simultaneous regeneration of an epidermis as well as a dermis. It was later appreciated that cell seeding speeds up epidermal regeneration but it appears otherwise not to be required for regeneration either of the epidermis or of the dermis. The evidence cited above led to identification of a cellfree macromolecular network with highly specific structure, the cell-free skin regeneration template (SRT), which alone possessed this unprecedented morphogenetic activity. Only one of several collagen-GAG matrices studied as described above was capable of preventing scar tissue formation and promoting dermal regeneration. The active ECM analog was characterized by a collagen/GAG ratio of 98/2 w/w, average pore diameter between 20 and 120 µm, random orientation of pore channels, and sufficiently high density of covalent crosslinks (average molecular weight between crosslinks in the template, 12 kDa) to resist degradation by collagenases for about 10 days following grafting. Several other very closely related ECM analogs showed either significantly reduced activity or no activity at all.9 In related studies it was observed that ECM analogs which showed high activity in promoting dermal regeneration also delayed significantly the onset of wound contraction.9 The available evidence compelled the conclusion that the activity of this insoluble network depended critically on maintenance of a highly specific three dimensional structure inside the wound bed over a period of time between about 5 and 15 days. The active network has been referred to as skin regeneration template (SRT). The observed activity of SRT, consisting in drastic modification of the outcome of the wound healing process, has not been duplicated by application on the wound bed of solutions of one or more growth factors or by application of suspensions of keratinocytes or fibroblasts. Regeneration of a partially functional sciatic nerve was induced across a transected gap of 15 mm in the rat sciatic nerve15-18 by another ECM analog, which was also found to possess a highly specific network structure. In this animal model the nerve stumps at either side of the gap were inserted in a silicone tube (tubulation); in the absence of a tube, regeneration was decidedly absent and neuroma formation was invariably reported. Although spontaneous regeneration of nerve through the tube occurred reproducibly at a gap length of 5 mm, regeneration across a 15 mm gap has not been observed.19,20 The study eventually led to identification of an ECM analog which, when filling the silicone tube and thereby connecting the two nerve stumps, induced regeneration across a 15 mm gap. This analog has been referred to as nerve regeneration template (NRT). It

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has an average pore diameter of 5 µm, an average molecular weight between crosslinks of 30-40 kDa, a preferred orientation of pore channel axes in the direction of the nerve axis, and a 98/2 w/w ratio of type I collagen to GAG.16,17 Clearly, there are significant differences between the network structure of skin and nerve regeneration templates. Yet a third ECM analog has been reported, capable of inducing regeneration of the canine meniscus following 80% transection.21,22 Although the ECM analog used in these studies has been stated by the authors to be similar to that described earlier,9 its detailed structure was not reported. There is, therefore, substantial evidence that, under appropriate conditions, the adult mammal can be induced to regrow certain organs which are not regenerated spontaneously. This exciting conclusion suggests new therapeutic avenues. However, there are still many unanswered questions about the detailed cell-biological mechanism by which active ECM analogs modify so spectacularly one or more of the processes involved in conventional wound healing.

Methodological Principles of Induced Regeneration Over the years it has become clear that a number of conditions must be in place in order to conduct a reproducible study of induced regeneration. One of the most important is the choice of an anatomically well-defined lesion (experimental volume) in which to study the phenomena. This concept was pioneered in the studies of Billingham and Medawar1 as a means of minimizing animal to animal variability in observations. The experimental volume can be isolated by containing it within anatomically distinct tissues belonging to a neighboring organ or by a device or (in the case of skin) by the atmosphere. Since mammals possess a very small but finite potential for regeneration, it is necessary to “correct” the observed regeneration by subtracting the amount which has been contributed by spontaneous regeneration. This requirement is equivalent to defining a negative control in every experimental model, consisting of the observations obtained in the absence of the presumptive regeneraion template. Although the mechanistic interpretation of induced regeneration has not been completed, two arguments have been advanced to support the contention that there is a requirement for structuring a regeneration template as an ECM analog. First, studies of organ development have shown conclusively that the early presence of specific ECM components is required for formation of physiological organs.23-25 In contrast, the exudate which flows into a spontaneously healing wound after injury does not contain components of the insoluble and nondiffusible ECM networks (although it contains several types of cells and regulators); the healing process in this case eventually leads to synthesis of scar. Second, the evidence presented in the preceding section clearly supports the contention that specific ECM analogs indeed suffice to lead to regeneration of at least two organs. However, there is clearly insufficient evidence to conclude definitively that regeneration can be induced only by specific ECM analogs.

In Vivo Synthesis of Organs Using Collagen-GAG Copolymers

Conclusive evidence that an ECM analog has induced organ regeneration can only be obtained with a performance assay which simply compares the structure of the regenerate to that of a normal control. Assays of this type can be destructive (e.g., histology), in which case only one experimental time point can be obtained per animal, or nondestructive, such that the complete kinetic curve for regeneration can be obtained from a single animal. The fidelity of regeneration is a figure of merit which is assigned to each candidate template. The regeneration template is then the ECM analog which has scored a very high value of the fidelity index.

Models of the Mechanism of Regeneration I summarize below the major physicochemical models26-28 which have been used to compile a mechanistic interpretation of the process of regeneration. It is evident that a regeneration template modifies drastically the kinetics and mechanism of spontaneous wound healing which normally lead to wound contraction and scar synthesis. However, a view of a template as a classical catalyst is not supported by the well known fact that regeneration templates are necessarily consumed during the process which they modify. Micrometer Scale Proximity of Host to Template Surface Cells and cytokines present in the exudate are transferred to the surface of the template following implantation of the porous template into the wound bed. Surface tension forces pull the exudate inside the capillaries (pore channels) of the template, as described by: P = 2γ/r

(1)

where r is the radius of the pore channel in a template undergoing wetting by exudate with an air-liquid surface tension of γ in dyne/cm. Wetting of the opposed surfaces is promoted as the suction pressure P in Equation (1) increases, and such increase is occasioned by a decrease in average pore radius. The values reached by P are not negligible; water with an air-liquid surface tension of γ = 72 dyne/cm, for example, is pulled inside a pore radius of 100 µm with a suction pressure of almost one-hundredth of one atmosphere and P increases almost to one full atmosphere when the pore radius decreases to a value as low as 1 µm. Under conditions where exudate flows inside the pore channels of a template with average pore diameter of 100 µm, cells and cytokines are within a distance of less than 50 µm from the template surface; this distance can then be covered by these components of the exudate within a few minutes. Maximum Dimension of Template, ~100 µm Cell migration from the solid-like tissue of the wound bed into the template requires the availability of adequate nutrition. The precise nutritional requirements of any cell are too complex to incorporate into a simple model; instead, these requirements will be simplified by defining a critical nutrient which is required for normal cell function; such a nutrient is assumed to be metabolized by the cell at a rate R

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mole/cm3/s. Picture the nutrient being transported from the solid-like tissue, where the concentration of nutrient is assumed to be a constant C0 due to the presence of vascular supply, over a distance L through the exudate, until it reaches the cell. Prior to the onset of angiogenesis, i.e., during the first few days following template implantation, the nutrient is transported exclusively by diffusion, characterized by a diffusivity D cm2/s. Dimensional analysis readily yields the cell lifeline number: S = RL2/DC0

(2)

which can be used to compare the relative magnitude of the rate of nutrient consumption by the cell nutrient (numerator) to the rate of supply of nutrient to the cell by diffusion (denominator). The cell dies if the rate of consumption of the critical nutrient exceeds greatly the rate of supply, S>>1. At steady state the rate of consumption of nutrient by the cell just equals the rate of transport by diffusion over the distance L. Under conditions of steady state, S = O(1); at that point, the value of L becomes the critical cell path length, Lc, the distance of migration beyond which cells require the presence of a vascular supply. Another way of looking at Lc is that it is the longest distance away from the wound bed boundary along which the cell can migrate without requiring nutrient in excess of that supplied by diffusion. For many cell nutrients of low molecular weight, Lc is of order 100 µm, an estimate of the maximum template dimension which can support cells. Template Specific Surface Is Bounded Following its migration onto the template surface a host cell is visualized interacting with binding sites on the surface. The surface density of binding sites can be expressed as Φb, equal by definition to the number of sites Nb per unit surface A of template. Φb can also be expressed more usefully in terms of quantities potentially measurable by optical microscopy, i.e., in terms of the volume density of binding sites ρb (number of sites per unit volume porous template) and the specific surface σ of the template expressed in units of mm2/cm3: Φb = Nb/A = ρb/σ

(3)

Assuming that each cell is bound to (an a priori unknown number of) χ binding sites, there will be Nb/χ bound cells per unit surface; the volume density of cells will be ρc = ρb/χ and the surface density of cells will be: Φc = Φb/χ = Nb/χA = ρb/χσ = ρc/σ

(4)

Levels of myofibroblast density inside templates with pore diameter of about 10 µm have been observed to reach typical values of the volume density, ρ c, of order 107 myofibroblasts per cm3 porous template. For a template of average pore diameter 10 µm the specific surface σ is calculated to be approximately 8 x 104 mm2/cm3 template; therefore, 1 cm3 porous template is characterized by a cell surface density of Φc = ρc/σ = 107/8 x 104 = 125 cells/mm2. For a

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template of identical composition but average pore diameter as large as 300 µm, Φc is the same as above; however, the specific surface is calculated to be only about 3 x 103 mm2/cm3 template. In this case, the volume density of cells is, accordingly, only ρc = Φc ⋅ σ = 125 x 3 x 103 = 3.75 x 105 per cm3 porous template. We conclude that the template which has the smaller average pore diameter (10 µm) has a volume density of myofibroblasts which is about 27 times lower than with the template which has the larger pore diameter (300 µm). The existence of a maximum pore diameter requirement for the template is suggested by these calculations, simply to ensure a specific surface which is large enough to bind an appropriately large number of cells. In addition, it is clear that cells originating in the wound bed cannot migrate inside the template and eventually reach binding sites on its surface unless the template has an average pore diameter large enough to allow for this. We conclude that there is, therefore, a requirement for a minimum pore diameter for the template, about equal to the characteristic diameter of the cells (of order 5 µm). Thus, the pore diameter of the regeneration template is limited both by an upper and a lower bound. This conclusion is in agreement with the experimental evidence which shows that ECM analogs, identical in chemical composition but differing only in average pore diameter, show maximum activity (inhibition of onset of wound contraction, consistent with regeneration rather than scar formation) when the average pore diameter lies between 20 and 120 µm.9 Further evidence has shown that, when other structural parameters of the template remain constant, loss of the 20-120 µm porous structure of the template by simple evaporation at room temperature (a process which yields an ECM analog with average pore diameter of less than 1 µm) leads to synthesis of a scar capsule at the surface of the grafted analog, evidence of a barrier to cell migration inside an implant.29 Critical Residence Time of Template The residence time of a template is also bounded. A template has to stay in place long enough to induce the appropriate synthetic processes to take place; soon after that, it must disappear in a timely fashion so as not to interfere with these same processes which it induces. The time period necessary to induce synthesis is roughly equal to that required to complete the wound healing process at that anatomical site. (In general, the rate of wound healing is quite different in tissues such as, say, the dermis and the sciatic nerve.) Since the template is an insoluble (and, therefore, nondiffusible) three dimensional network it follows that cells which are bound on it become immobilized and their migration is, accordingly, arrested. Not only cells are prevented from migrating to locations which are appropriate for synthesis of a new organ but, in addition, the laying down of newly synthesized ECM by the cells in the space of the wound bed is probably blocked physically by the presence of the template. It follows that the persisting insolubility of the template will increasingly interfere with the synthesis of the new organ at that site. To avoid this eventuality, the template is required to become diffusible (by degradation to small molecular fragments) and thereby disappear from the wound

Tissue Engineering of Prosthetic Vascular Grafts

bed in order not to impede cellular processes which lead to the emerging organ. A steady state model is probably the simplest way to accommodate these two requirements. It requires synchronization of the two processes: organ synthesis and template degradation.5,9 This model leads directly to the hypothesis of isomorphous tissue replacement: td/ts = O(1)

(5)

In Equation (5) td denotes a characteristic time constant for degradation of the template at the tissue site where a new organ is synthesized with a time constant of ts. The degradation rate can be estimated by histological observation of the decrease in mass of template fragments at various times.10,30 A closer estimate of td has been obtained by measuring the kinetics of disintegration of the macromolecular network using rubber elasticity theory.3 A third procedure consists of monitoring the kinetics of mass disappearance of a radioactively labeled template. A rough estimate of ts can be obtained by observing th, the timescale of synthesis of new tissue during healing (in the absence of a template) at the anatomical site.31 Using the latter approach it has been estimated that ts for the regenerating dermis is of order 3 weeks31 and of order 6 weeks for the regenerating peripheral nerve.16 These estimates allow adjustment of td for the template, by adjustment of the crosslink density and GAG content, to levels which are approximately equal to the values of ts, as the latter are dictated by the nature of the anatomical site. Experimental support for the isomorphous tissue replacement hypothesis has been based on observations such that, when the ratio in Equation (5) was adjusted to values much smaller than one (by implanting a rapidly degrading ECM analog, for which td