Natural-Based Polymers for Biomedical Applications

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Natural-based polymers for biomedical applications

© 2008, Woodhead Publishing Limited

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Related titles: Biomedical polymers (ISBN 978-1-84569-070-0) This book reviews the structure, processing and properties of biomedical polymers. It discusses the various groups of biopolymers including natural polymers and synthetic biodegradable and non-biodegradable polymers. Chapters also review the application of biomedical polymers in such areas as scaffolds for tissue engineering, drug delivery systems and cell encapsulation. The book also considers the use of polymers in replacement heart valves and arteries, in joint replacement and in biosensor applications. Tissue engineering using ceramics and polymers (ISBN 978-1-84569-176-9) Tissue engineering is a rapidly developing technique for the repair and regeneration of diseased tissue in the body. This authoritative and wide-ranging book reviews how ceramic and polymeric biomaterials are being used in tissue engineering. The first part reviews the nature of ceramics and polymers as biomaterials together with techniques for using them, such as building tissue scaffolds, transplantation techniques, surface modification and ways of combining tissue engineering with drug delivery and biosensor systems. The second part discusses the regeneration of particular types of tissue from bone and cardiac and intervertebral disc tissue to skin, liver, kidney and lung tissue. Surfaces and interfaces for biomaterials (ISBN 978-1-85573-930-7) This book presents our current level of understanding on the nature of a biomaterial surface, the adaptive response of the biomatrix to that surface, techniques used to modify biocompatibility, and state-of-the-art characterisation techniques to follow the interfacial events at that surface. Details of these and other Woodhead Publishing materials books, as well as materials books from Maney Publishing, can be obtained by: • visiting our web site at www.woodheadpublishing.com • contacting Customer Services (e-mail: [email protected]; fax: +44 (0) 1223 893694; tel: +44 (0) 1223 891358 ext: 130; address: Woodhead Publishing Ltd, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, England) If you would like to receive information on forthcoming titles, please send your address details to: Francis Dodds (address, tel. and fax as above; e-mail: [email protected]). Please confirm which subject areas you are interested in. Maney currently publishes 16 peer-reviewed materials science and engineering journals. For further information visit www.maney.co.uk/journals

© 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications Editor-in-Chief: Rui L. Reis Section Editors: Nuno M. Neves, João F. Mano, Manuela E. Gomes, Alexandra P. Marques and Helena S. Azevedo

Woodhead Publishing and Maney Publishing on behalf of The Institute of Materials, Minerals & Mining WPTF2005

CRC Press Boca Raton Boston New York Washington, DC

WOODHEAD

PUBLISHING LIMITED

Cambridge England

© 2008, Woodhead Publishing Limited

iv Woodhead Publishing Limited and Maney Publishing Limited on behalf of The Institute of Materials, Minerals & Mining Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington Cambridge CB21 6AH, England www.woodheadpublishing.com Published in North America by CRC Press LLC, 6000 Broken Sound Parkway, NW, Suite 300, Boca Raton, FL 33487, USA First published 2008, Woodhead Publishing Limited and CRC Press LLC © 2008, Woodhead Publishing Limited The authors have asserted their moral rights. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from Woodhead Publishing Limited. The consent of Woodhead Publishing Limited does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited for such copying. Trademark notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Cataloging in Publication Data A catalog record for this book is available from the Library of Congress. Woodhead Publishing ISBN 978-1-84569-264-3 (book) Woodhead Publishing ISBN 978-1-84569-481-4 (e-book) CRC Press ISBN 978-1-4200-7607-3 CRC Press order number WP7607 The publishers’ policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp which is processed using acid-free and elementary chlorine-free practices. Furthermore, the publishers ensure that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by Replika Press Pvt Ltd, India Printed by T J International Limited, Padstow, Cornwall, England

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Contents

Contributor contact details

xvii

Preface

xxiii

Part I Sources, properties, modification and processing of natural-based polymers 1

Polysaccharides as carriers of bioactive agents for medical applications

3

R. PAWAR, W. JADHAV, S. BHUSARE and R. BORADE, Dnyanopasak College, India, S. FARBER, D. ITZKOWITZ and A. DOMB, The Hebrew University, Jerusalem, Israel

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 1.10 1.11 1.12 1.13 1.14 1.15 1.16 1.17 1.18 1.19

Introduction Starch Cellulose Heparinoid (sulfated polysaccharides) Dextran Pectin Arabinogalactan Drug conjugated polysaccharides Polysaccharide dextrans Mannan Pullulan Polysaccharide macromolecule–protein conjugates Cationic polysaccharides for gene delivery Diethylaminoethyl-dextran Polysaccharide–oligoamine based conjugates Chitosan Applications of polysaccharides as drug carriers Applications of dextran conjugates Site-specific drug delivery

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3 6 7 8 10 12 13 15 19 22 23 24 25 26 27 27 31 33 38

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Contents

1.20 1.21 1.22

Pectin drug site-specific delivery Liposomal drug delivery References

38 40 45

2

Purification of naturally occurring biomaterials

54

M. N. GUPTA, Indian Institute of Technology Delhi, India

2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8 2.9 2.10 2.11 2.12 3

Introduction Classes of naturally occurring biomaterials Downstream processing of small molecular weight natural products Purification strategies for proteins Purification of lipids Purification of polysaccharides Purification of nucleic acids Purification of complex biomaterials Future trends Acknowledgement Sources of further information References

54 55 57 60 67 71 72 75 76 77 77 78

Processing of starch-based blends for biomedical applications

85

R. A. SOUSA, V. M. CORRELO, S. CHUNG, N. M. NEVES, J. F. MANO and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

3.1 3.2 3.3 3.4 3.5

Introduction Starch Starch-based blends Conclusions References

4

Controlling the degradation of natural polymers for biomedical applications

85 85 88 98 99 106

H. S. AZEVEDO, T. C. SANTOS and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

4.1 4.2 4.3 4.4 4.5

Introduction The importance of biodegradability of natural polymers in biomedical applications Degradation mechanisms of natural polymers and metabolic pathways for their disposal in the body Assessing the in vitro and in vivo biodegradability of natural polymers Controlling the degradation rate of natural polymers

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Contents

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4.6 4.7 4.8

Concluding remarks Acknowledgements References

124 125 125

5

Smart systems based on polysaccharides

129

M. N. GUPTA and S. RAGHAVA, Indian Institute of Technology Delhi, India

5.1 5.2 5.3 5.4 5.5 5.6 5.7 5.8 5.9 5.10

What are smart materials? Chitin and chitosan Alginates Carrageenans Other miscellaneous smart polysaccharides and their applications Polysaccharide-based composite materials Future trends Acknowledgement Sources of further information References

129 131 136 140 145 146 149 152 152 154

Part II Surface modification and biomimetic coatings 6

Surface modification for natural-based biomedical polymers

165

I. PASHKULEVA, P. M. LÓPEZ-PÉREZ and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

6.1 6.2 6.3 6.4 6.5 6.6 6.7 6.8 6.9

Introduction Some terms and classifications Wet chemistry in surface modification Physical methods for surface alterations Grafting Bio-approaches: Mimicking the cell–cell interactions Future trends Acknowledgements References

165 165 167 171 177 179 186 186 186

7

New biomineralization strategies for the use of natural-based polymeric materials in bone-tissue engineering

193

I. B. LEONOR, S. GOMES, P. C. BESSA, J. F. MANO, R. L. REIS, 3B’s Research Group, University of Minho, Portugal and M. Casal, CBMA – Molecular and Environmental Biology Center, University of Minho, Portugal

7.1

Introduction

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Contents

7.2 7.3 7.4 7.5 7.6 7.7

The structure, development and mineralization of bone Bone morphogenetic proteins in tissue engineering Bio-inspired calcium-phosphate mineralization from solution General remarks and future trends Acknowledgments References

194 201 206 216 217 217

8

Natural-based multilayer films for biomedical applications

231

C. PICART, Université Montpellier, France

8.1 8.2 8.3 8.4 8.5 8.6 8.7 8.8 9

Introduction Physico-chemical properties Different types of natural-based multilayer films for different applications Bioactivity, cell adhesion, and biodegradability properties Modulation of film mechanical properties Future trends Sources of further information and advice References Peptide modification of polysaccharide scaffolds for targeted cell signaling

231 234 240 244 248 250 251 252

260

S. LÉVESQUE, R. WYLIE, Y. AIZAWA and M. SHOICHET, University of Toronto, Canada

9.1 9.2 9.3 9.4 9.5 9.6 9.7

Introduction Polysaccharide scaffolds in tissue engineering Peptide immobilization Measurement Challenges associated with peptide immobilization Tissue engineering approaches targeting cell signaling References

260 265 267 272 274 275 277

Part III Biodegradable scaffolds for tissue regeneration 10

Scaffolds based on hyaluronan derivatives in biomedical applications

291

E. TOGNANA, Fidia Advanced Biopolymers s.r.l., Italy

10.1 10.2 10.3 10.4

Introduction Hyaluronan Hyaluronan-based scaffolds for biomedical applications Clinical applications

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291 291 293 298

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10.5 10.6 10.7

Future trends Sources of further information and advice References

308 309 310

11

Electrospun elastin and collagen nanofibers and their application as biomaterials

315

R. SALLACH and E. CHAIKOF, Emory University/Georgia Institute of Technology, USA

11.1 11.2 11.3 11.4 11.5 11.6 11.7 11.8 11.9 12

Introduction Electrospinning as a biomedical fabrication technology Generation of nanofibers with controlled structures and morphology Generation of collagen and elastin small-diameter fibers and fiber networks Biological role of elastin Generation of crosslinked fibers and fiber networks Multicomponent electrospun assemblies Future trends References

315 316

318 321 328 329 331 332

Starch-polycaprolactone based scaffolds in bone and cartilage tissue engineering approaches

337

317

M. E. GOMES, J. T. OLIVEIRA, M. T. RODRIGUES, M. I. SANTOS, K. TUZLAKOGLU, C. A. VIEGAS, I. R. DIAS and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

12.1 12.2 12.3

337 338

12.7 12.8 12.9

Introduction Starch+ ε-polycaprolactone (SPCL) fiber meshes SPCL-based scaffold architecture, stem cell proliferation and differentiation In vivo functionality of SPCL fiber-mesh scaffolds Cartilage tissue engineering using SPCL fiber-mesh scaffolds Advanced approaches using SPCL scaffolds for bone tissue engineering aiming at improved vascularization Conclusions Acknowledgments References

13

Chitosan-based scaffolds in orthopedic applications

357

12.4 12.5 12.6

339 341 342 346 350 351 351

K. TUZLAKOGLU and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

13.1

Introduction: Chemical and physical structure of chitosan and its derivatives

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Contents

13.2 13.3 13.4 13.5 13.6

Production methods for scaffolds based on chitosan and its composites or blends Orthopedic applications Conclusions and future trends Acknowledgements References

358 365 369 369 369

14

Elastin-like systems for tissue engineering

374

J. RODRIGUEZ-CABELLO, A. RIBEIRO, J. REGUERA, A. GIROTTI and A. TESTERA, Universidad de Valladolid, Spain

14.1 14.2 14.3

14.13 14.14 14.15

Introduction Genetic engineering of protein-based polymers Genetic strategies for synthesis of protein-based polymers State-of-the-art in genetically-engineered protein-based polymers (GEBPs) Elastin-like polymers Self-assembly behaviour of peptides and proteins Self-assembly of elastin-like polymers (ELPs) Biocompatibility of ELPs Biomedical applications ELPs for drug delivery Tissue engineering Self-assembling properties of ELPs for tissue engineering Processability of ELPs for tissue engineering Future trends References

15

Collagen-based scaffolds for tissue engineering

14.4 14.5 14.6 14.7 14.8 14.9 14.10 14.11 14.12

374 375 376 377 377 379 379 381 382 382 383 388 388 389 391 396

G. CHEN, N. KAWAZOE and T. TATEISHI, National Institute for Materials Science, Japan

15.1 15.2 15.3 15.4 15.5 15.6 15.7 15.8 15.9

Introduction Structure and properties of collagen Collagen sponge Collagen gel Collagen–glycosoaminoglycan (GAG) scaffolds Acellularized scaffolds Hybrid scaffolds Future trends References

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396 396 397 400 402 404 405 409 409

Contents

16

Polyhydroxyalkanoate and its potential for biomedical applications

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416

P. FURRER and M. ZINN, Swiss Federal Laboratories for Materials Testing and Research (Empa), Switzerland, and S. PANKE, Swiss Federal Institute of Technology (ETH), Switzerland

16.1 16.2 16.3 16.4 16.5 16.6 16.7

Introduction Biosynthesis Chemical digestion of non-PHA biomass Purification of PHA Potential applications of PHA in medicine and pharmacy Conclusions and future trends References

416 417 425 431 434 437 437

17

Electrospinning of natural proteins for tissue engineering scaffolding

446

P. I. LELKES, M. LI, A. PERETS, L. LIN, J. HAN and D. WOERDEMAN, Drexel University, USA

17.1 17.2 17.3 17.4 17.5 17.6 17.7 17.8 17.9

Introduction The electrospinning process Electrospinning natural animal polymers Electrospinning blends of synthetic and natural polymers Electrospinning novel natural ‘green’ plant polymers for tissue engineering Cellular responses to electrospun scaffolds: Does fiber diameter matter? Conclusions and future trends Sources of further information and advice References

446 448 455 460 466 474 474 475 476

Part IV Naturally-derived hydrogels: Fundamentals, challenges and applications in tissue engineering and regenerative medicine 18

Hydrogels from polysaccharide-based materials: Fundamentals and applications in regenerative medicine

485

J. T. OLIVEIRA and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

18.1 18.2

Introduction: Definitions and properties of hydrogels Applications of hydrogels produced from different polysaccharides in tissue engineering and regenerative medicine

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487

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Contents

18.3 18.4 18.5 18.6 18.7 18.8 18.9 18.10 18.11 18.12 18.13 18.14 18.15

Agarose Alginate Carrageenan Cellulose Chitin/chitosan Chondroitin sulfate Dextran Gellan Hyaluronic acid Starch Xanthan Conclusion References

488 489 491 492 493 495 496 497 498 500 501 502 503

19

Alginate hydrogels as matrices for tissue engineering

515

H. PARK and K.-Y. LEE, Hanyang University, South Korea

19.1 19.2 19.3 19.4 19.5 19.6

Introduction Properties of alginate Methods of gelling Applications of alginate hydrogels in tissue engineering Summary and future trends References

515 516 520 523 528 528

20

Fibrin matrices in tissue engineering

533

B. TAWIL, H. DUONG and B. WU, University of California Los Angeles, USA

20.1 20.2 20.3 20.4 20.5 20.6 20.7 20.8 20.9

Introduction Fibrin formation Fibrin use in surgery Fibrin matrices to deliver bioactive molecules Fibrin – cell constructs Mechanical characteristics of fibrin scaffold Future trends Conclusions References

533 534 535 535 536 540 541 542 543

21

Natural-based polymers for encapsulation of living cells: Fundamentals, applications and challenges

549

P. DE VOS, University Hospital of Groningen, The Netherlands

21.1

Introduction

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Contents

21.2

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21.3 21.4 21.5 21.6 21.7

Approaches of encapsulation: Materials and biocompatibility issues Physico-chemistry of microcapsules and biocompatibility Immunological considerations Conclusions and future trends Sources of further information and advice References

550 556 559 561 563 564

22

Hydrogels for spinal cord injury regeneration

570

A. J. SALGADO and N. SOUSA, Life and Health Sciences Research Institute (ICVS), University of Minho, Portugal, and N. A. SILVA, N. M. NEVES and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

22.1 22.2 22.3 22.4 22.5 22.6 22.7

Introduction Brief insights on central nervous system biology Current approaches for SCI repair Hydrogel-based systems in SCI regenerative medicine Conclusions and future trends Acknowledgements References

570 571 576 578 587 588 588

Part V Systems for the sustained release of molecules 23

Particles for controlled drug delivery

597

E. T. BARAN and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

23.1 23.2 23.3 23.4 23.5 23.6 23.7 23.8 23.9

Introduction Novel particle processing methods Hiding particles: The stealth principle Finding the target Delivery of bioactive agents at the target site and novel deliveries Viral delivery systems Conclusions Acknowledgements References

597 597 602 604 608 611 612 613 613

24

Thiolated chitosans in non-invasive drug delivery

624

A. BERNKOP-SCHNÜRCH, Leopold-Franzens University, Austria

24.1 24.2

Introduction Thiolated chitosans

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624 625

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Contents

24.3 24.4 24.5 24.6 24.7

Properties of thiolated chitosans Drug delivery systems In vivo performance Conclusion References

625 633 634 638 639

25

Chitosan–polysaccharide blended nanoparticles for controlled drug delivery

644

J. M. ALONSO and F. M. GOYCOOLEA, Universidad de Santiagó de Compostela, Spain, and I. HIGUERA-CIAPARA, Centro de Investigación en Alimentación y Desarrollo, Mexico

25.1 25.2 25.3 25.4 25.5 25.6 25.7 25.8 25.9

Introduction Polysaccharides in nanoparticle formation Nanoparticles constituted by chitosan Drug delivery properties and biopharmaceutical applications Hybrid nanoparticles consisting of chitosan and other polysaccharides Future trends Sources of further information and advice Acknowledgements References

644 645 651 654 656 668 668 671 671

Part VI Biocompatibility of natural-based polymers 26

In vivo tissue responses to natural-origin biomaterials

683

T. C. SANTOS, A. P. MARQUES and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

26.1 26.2 26.3 26.4

26.6 26.7 26.8

Introduction Inflammation and foreign-body reactions to biomaterials Role of host tissues in biomaterials implantation Assessing the in vivo tissue responses to natural-origin biomaterials Controlling the in vivo tissue reactions to natural-origin biomaterials Final remarks Acknowledgements References

693 695 695 695

27

Immunological issues in tissue engineering

699

26.5

683 684 686 690

N. ROTTER, Ulm University, Germany

27.1

Introduction

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Contents

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27.2 27.3 27.4 27.5 27.6 27.7 27.8 27.9 27.10

Immune reactions to biomaterials Host reactions related to the implant site Immune reactions to different types of cells Immune reactions to in vitro engineered tissues Immune protection of engineered constructs Strategies directed towards reactions to biomaterials Strategies directed towards reactions to implanted cells Future trends References

699 701 701 704 705 706 707 709 710

28

Biocompatibility of hyaluronic acid: From cell recognition to therapeutic applications

716

K. GHOSH, Children’s Hospital and Harvard Medical School, USA

28.1 28.2 28.3 28.4 28.5 28.6 28.7 28.8

Introduction Native hyaluronan Therapeutic implications of native hyaluronan Engineered hyaluronan Implications for regenerative medicine Conclusion Future trends References

716 717 721 722 727 728 728 728

29

Biocompatibility of starch-based polymers

738

A. P. MARQUES, R. P. PIRRACO and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

29.1 29.2 29.3 29.4 29.5 29.6 29.7

Introduction Starch-based polymers in the biomedical field Cytocompatibility of starch-based polymers Immunocompatibility of starch-based polymers Conclusions Acknowledgements References

738 740 745 748 752 753 753

30

Vascularization strategies in tissue engineering

761

M. I. SANTOS, and R. L. REIS, 3B’s Research Group, University of Minho, Portugal

30.1 30.2 30.3 30.4

Introduction Biology of vascular networks – angiogenesis versus vasculogenesis Vascularization: The hurdle of tissue engineering Neovascularization of engineered bone

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761 761 762 763

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Contents

30.5

Strategies to enhance vascularization in engineered grafts In vivo models to evaluate angiogenesis in tissue engineered products Future prospects Sources of further information and advice References

30.6 30.7 30.8 30.9

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765 774 776 776 776

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Contributor contact details

(* = main contact)

Editors

Chapter 2

Rui L. Reis,* Nuno M. Neves, João F. Mano, Manuela E. Gomes, Alexandra P. Marques, Helena S. Azevedo 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

M. N. Gupta Department of Chemistry Indian Institute of Technology Delhi Hauz Khas New Delhi 110 016 India

E-mail: [email protected] [email protected] [email protected] [email protected] [email protected] [email protected]

Chapter 1 A. J. Domb Department of Medicinal Chemistry and Natural Products School of Pharmacy – Faculty of Medicine The Hebrew University Jerusalem Israel E-mail: [email protected] © 2008, Woodhead Publishing Limited

E-mail: [email protected]

Chapter 3 R. A. Sousa 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal E-mail: [email protected]

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Contributor contact details

Chapter 4

Chapter 7

H. S. Azevedo 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

I. B. Leonor 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

E-mail: [email protected]

E-mail: [email protected]

Chapter 5

Chapter 8

M. N. Gupta Department of Chemistry Indian Institute of Technology Delhi Hauz Khas New Delhi 110 016 India

C. Picart DIMNP Dynamique des Interactions Membranaires Normales et Pathologiques CNRS UMR5235 Université Montpellier II et I cc 107 34 095 Montpellier France

E-mail: [email protected]

Chapter 6 I. Pashkuleva 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal E-mail: [email protected]

E-mail: [email protected]

Chapter 9 M. S. Shoichet Terrence Donnelly Centre for Cellular and Biomolecular Research University of Toronto 160 College Street Room 514 Toronto Ontario M5S 3E1 Canada E-mail: [email protected]

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Contributor contact details

xix

Chapter 10

Chapter 13

E. Tognana R&D – Head of Unit Fidia Advanced Biopolymers s.r.l. Via Ponte della Fabbrica 3\b 35031 Abano Terme PD Italy

K. Tuzlakoglu 3B’s Research Group Biomaterials, Biodegradables and Biomimetics Dept. of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

E-mail: [email protected]

Chapter 11 R. E. Sallach Emory University 101 Woodruff Circle Room 5105 Atlanta GA 30322 USA E-mail: [email protected]

Chapter 12 M. E. Gomes 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

E-mail: [email protected]

Chapter 14 J. Rodriguez-Cabello G. I. R. Bioforge Dept. Física de la Materia Condensada Universidad de Valladolid Spain E-mail: [email protected] [email protected]

Chapter 15 G. Chen Biomaterials Center National Institute for Materials Science 1-1 Namiki Tsukuba Ibaraki 305-0044 Japan

E-mail: [email protected] E-mail: [email protected]

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Contributor contact details

Chapter 16

Chapter 19

M. Zinn Laboratory for Biomaterials Swiss Federal Laboratories for Materials Testing and Research (Empa) Lerchenfeldstrasse 5 CH-9014 St. Gallen Switzerland

K.-Y. Lee Department of Bioengineering Hanyang University 17 Haengdang-dong Seongdong-gu Seoul 133-791 South Korea Email: [email protected]

E-mail: [email protected]

Chapter 20 Chapter 17 P. I. Lelkes Drexel University Laboratory of Cellular Tissue Engineering School of Biomedical Engineering Science and Health Systems Bossone 707 3141 Chestnut Street Philadelphia PA 19104 USA E-mail: [email protected]

Chapter 18 J. T. Oliveira 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal E-mail: [email protected]

© 2008, Woodhead Publishing Limited

B. Tawil University of California Los Angeles Department of Bioengineering 7523 Boelter Hall Los Angeles CA 90095-1600 USA E-mail: [email protected]

Chapter 21 P. De Vos Department of Pathology and Laboratory Medicine University Hospital of Groningen Hanzeplein 1 9700 RB Groningen The Netherlands E-mail: [email protected]

Contributor contact details

xxi

Chapter 22

Chapter 25

A. Salgado Life and Health Sciences Research Institute (ICVS) School of Health Sciences University of Minho Campus de Gualtar 4710-057 Braga Portugal

M. Alonoso Faculty of Pharmacy Universidad de Santiago de Compostela Campus Universitario Sur s/n 15782 Santiago de Compostela A Coruña Spain

E-mail: [email protected]

E-mail: [email protected]

Chapter 23

Chapter 26

E. T. Baran 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

T. C. Santos 3 B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

E-mail: [email protected]

Email: [email protected]

Chapter 24

Chapter 27

A. Bernkop-Schnürch Institute of Pharmacy Leopold-Franzens University Innsbruck Innrain 52 Josef Möller Haus 6020 Innsbruck Austria

N. Rotter Department of Otorhinolaryngology Ulm University Frauensteige 12 89075 Ulm Germany

E-mail: [email protected]

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E-mail: [email protected]

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Contributor contact details

Chapter 28

Chapter 30

K. Ghosh Karp Family Research Laboratory Room 11.005E Vascular Biology Program Children’s Hospital and Harvard Medical School 300 Longwood Avenue Boston MA 02115 USA

M. I. Santos 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics Department of Polymer Engineering University of Minho Campus de Gualtar 4710-057 Braga Portugal

E-mail: [email protected]

E-mail: [email protected]

Chapter 29 A. P. Marques 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics University of Minho Campus de Gualtar 4710-057 Braga Portugal E-mail: [email protected]

© 2008, Woodhead Publishing Limited

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Preface

Polymers and polymeric-based systems play a key role in most devices used in distinct biomedical applications. Among them, polymers of natural origin are one of the most attractive options, mainly due to their similarities with the extracellular matrix and other polymers found in the human body. Such systems are also chemically versatile, may be modified by well established chemical methods and usually exhibit a rather good biological performance. This book describes both the most widely studied, as well as some of the most promising naturally-derived polymers that have been more recently suggested for use as implantable biomaterials, as controlled release carriers or scaffolds for tissue engineering. The organization of the different sections aims to provide the reader with a comprehensive overview of the most important topics covering the use of natural-based polymers in the biomedical area. The book is aimed to be used by a wide variety of readers working in industry and academia, as well as undergraduate and postgraduate students. Part I is dedicated to a detailed review of sources and properties of naturalbased polymers for biomedical applications. The section includes an indepth analysis of polysaccharide biomaterials and their derivatives. The properties of this class of materials are extensively reviewed, as well as strategies to modify them for specific applications, in particular tissue engineering and controlled release devices. This part of the book also covers aspects of processing of natural-based materials, the key issue of the control of the kinetics of degradation and, finally, opportunities and strategies to design smart systems exploring the specific properties of natural-based polymers. As biomaterials are in contact with tissues or body fluids, the surface plays an important role in the performance and biocompatibility of medical devices. Part II is devoted to the description of how surfaces of biomaterials based on natural-based polymers may be modified through a variety of methodologies and how this could influence their biological behaviour. Examples are physicochemical routes that will change parameters such as chemical energy and roughness or biomimetic coatings which are especially relevant for orthopaedic applications. Moreover, nanotechnologies, and in

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Preface

particular the layer-by-layer technique, may be used to coat and modify surfaces with natural-based polyelectrolytes, providing a suitable method to control surface properties and to provide functional characteristics to the surface. Many peptides promote cell attachment in a specific manner because the motif is recognized by adhesion cell-membrane receptors such as integrins. As explored in Chapter 6, a suitable method to promote desirable cellular function is to modify materials such as polysaccharides with peptide motifs that could induce cell adhesion onto their surfaces. Part III is dedicated to a description of a number of natural-based biodegradable polymers prepared by different methodologies into scaffolds and/or hydrogels with specific application in tissue engineering and regeneration. The various chapters included in this section are mainly focused on protein and polysaccharide-based materials which present promising characteristics to be used as support materials for the regeneration of a variety of soft and/or hard tissues. The unique properties of natural polymers, such as pseudoplastic behaviour, gelation ability, water binding capacity, biodegradability and similarity to the extracellular matrix, make them indispensable partners in bioencapsulation technology. Part IV describes the use of natural-gelling polymers as matrix environments for the encapsulation of different therapeutic agents, including proteins, stem cells or genetically engineered cells, and their applications in tissue engineering and regenerative medicine. Controlled/sustained drug (bioactive agents) delivery systems have attracted much attention over the years due to their great importance in human medicine. The criteria for choosing the materials to act as carriers are challenging. Natural polymers have been considered in the design of novel drug delivery systems because they can be easily modified and processed into adequate matrices (such as nano/microparticles and hydrogels) for the effective delivery of bioactive agents. Part V focuses on delivery systems based on naturalorigin polymers for the controlled release of small molecular weight drugs and more unstable macromolecules such as hormones, enzymes and growth factors. The different chapters of this section provide very comprehensive reviews that deal with the technological challenges and emerging research needed to develop advanced drug delivery systems for therapeutic use. Part VI discusses the translation of the properties of natural-origin materials into their effective biological performance in varied biomedical applications. Several chapters highlight specific properties that are considered critical for a specific cell response, thus demonstrating the potential of natural-origin materials in tissue regeneration. Chapters provide a clear overview of hosttransplant reactions triggered by the implantation of natural-origin biomaterials and strategies either to prevent or benefit from these reactions in the context of tissue engineering. To our knowledge this is the most comprehensive and up-to-date book in

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xxv

the field of natural origin based biomedical polymers. For those of us that have been working for a long time in this demanding but intellectuallyrewarding area of research, it was a pleasure to prepare this book. We thank all the authors for their well-prepared and authoritative contributions. We hope this is a useful book and that you enjoy reading it as much as we did preparing it! Rui L. Reis Nuno M. Neves João F. Mano Manuela E. Gomes Alexandra P. Marques Helena S. Azevedo

© 2008, Woodhead Publishing Limited

Part I Sources, properties, modification and processing of natural-based polymers

1 © 2008, Woodhead Publishing Limited

1 Polysaccharides as carriers of bioactive agents for medical applications R. P A WA R, W. J A D H AV, S. B H U S A R E and R. B O R A D E, Dnyanopasak College, India, S. F A R B E R, D. I T Z K O W I T Z and A. D O M B, The Hebrew University, Jerusalem, Israel

1.1

Introduction

Carbohydrates occur in nature in the form of polysaccharides of medium or high molecular weights. These macromolecules are used in building the fundamental components of life. They serve mainly two functions: as energy yielding fuel and extra cellular structural elements. Polysaccharides are made either of one type of small unit, or with two alternating units that are not only informational molecules such as protein and nucleic acids. However, small polymers of six or more different unit of sugars connected in a branch chain shows different structures and stereochemistry which give information about their recognition in comparison with other macromolecules. The most abundant polysaccharides in nature are starch and cellulose made up of repeating Dglucose molecules. They are also known as glycans, which differ from each other in their monosaccharide units; in the length of a chain; in the types of the linking units and in the degree of branching. Polysaccharides are of mainly two types. Those that are made up of one type of monomer unit are called homopolysaccharides, while polysaccharides made up of two or more types of monomer unit are called heteropolysaccharides. Most homopolysaccharides serve as storage of fuels, such as starch and glycogens. Cellulose and chitin serve as structural elements in plant cell walls and animal exoskeletons. Heteropolysaccharides provide an extracellular support for microorganisms such as bacteria and animal tissue. An extracellular space is occupied by different heteropolysaccharides, which forms a matrix that holds individual cells together and gives protection, shape and support to the cells, tissues and organs. Polysaccharides do not have specific molecular weights. This difference in molecular weight is due to the cosequence of the mechanism of two types of polymer formation. The polysaccharide synthesis is a natural process of polymerization of monomeric units catalyzed by certain enzymes.1 The classification of polysaccharides is on the basis of the monosaccharide components present and the sequences of linkages between them. The 3 © 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications

classification of polysaccharides also depends on the anomeric configuration of linkages, ring size (furanose or pyranose), absolute configuration (D- or L-) and other substituents present. Polysaccharides with nucleic acids and proteins determine the functionality and specificity of the species.2 The physicochemical properties of polysaccharides depends on certain structural characteristics such as chain conformation and intermolecular associations. The regular order of polysaccharides has been used to assume a limited number of conformations because of severe steric restrictions on the freedom of rotation of sugar molecules interunit glycosidic bonds. The most stable arrangement of atoms in a polysaccharide molecule is that which satisfies the intra- and inter- molecular forces. The structural non-starch polysaccharides cellulose and xylan, have preferred orientations that automatically support extended conformations. The chains in some polysaccharides amylopectin tend to form wide helical conformations. The regularity and the degree of stiffness of polysaccharide chains affect the rate of fermentation. Pentose sugars such as arabinose and xylose adopt one conformation out of two: furanose rings (arabinose) that can oscillate and are more flexible, and pyranose rings (xylose and glucose) which are less flexible. Pectins, made up of galacturonic acid residues forms more flexible extended conformations possessing regular ‘hairy’ regions with pendant arabinogalactans. Carbohydrates, containing large numbers of hydroxyl groups are not only hydrophilic but they are also capable of generating apolar surfaces depending on the monomer ring conformation, the epimeric structure, and the stereochemistry of the glycosidic linkages. Apolarity of dextrin, glucans, and cellulose results in the decrease in the hydrophobic nature in solution. Hydrophobicity will also be affected by the degree of polysaccharide hydration, i.e. the greater hydrophobic nature of polysaccharides decreases their interaction with water. Carbohydrates contain several hydroxyl groups that interact with two water molecules each if they are not interacting with other hydroxyl groups on the molecule. The interaction between hydroxyl groups on the same or neighboring polysaccharides reduces their hydration status. β-linkages present at 3- and 4-positions in mannose or glucose homopolymers allow strong inflexible inter-residue hydrogen bonding, which reduces the hydration of polymer and gives rise to a rigid inflexible polysaccharide structure, whereas α-linkages present at 2-, 3- and 4-positions in mannose or glucose homopolymers increase the hydration of polymer for more flexible linkages.3 Since the begining of the 1990s, glycoscience (polysaccharides) became popular worldwide due to their involvement in several medicinal applications and performed a wide range of biological functions. 4 The sulfated polysaccharide, heparin, plays an important role in blood coagulation.5 Polysaccharide hyaluronan, acts as a lubricant for human joints. It also has been used to protect the corneal endothelium during ophthalmologic surgery.6 In addition, hyaluronan is not only used for lubricating and cushioning

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Polysaccharides as carriers of bioactive agents

5

properties, but is also used as an antiinflammatory agent along with chondroitin sulfate for the treatment of osteoarthritis and rheumatoid arthritis.7 Cell surface polysaccharides are used for numerous biological functions, including the recognition of cell, adhesion, regulation in cell growth, cancer metastasis, and inflammation.8 Polysaccharides are also used as attachment sites for infectious bacteria, viruses, toxins and hormones, which may may result in pathogenesis.9 Cell surface glycopolymers are heterogeneous polysaccharides with well defined chemical structures. Synthetic polysaccharide derivatives are important tools for the determination of carbohydrate based interactions.10–13 These synthetic polysaccharides with pendant sugar residues are not only simplified models of biopolymers bearing oligosaccharides but also work as artificial glycoconjugates in biochemistry and medicine. They are used as surfactants,14 texture-enhancing food additives,15 reverse osmosis membranes, and biologically active polymers.16–17 Sulfonated dextran and pentosan possess anticoagulant activities such as heparin.18 Polysaccharide based vaccines such as tumor-associated carbohydrate antigens (e.g. sTn), are effective against tumors.19 Different synthetic polysaccharides are biocompatible and biodegradable and are used in tissue engineering and controlled drug release devices. N-(2-hydroxypropyl) methacrylamide copolymers modified with galactosamine interact with asialoglycoprotein receptors on hepatocytes and hepatocarcinomas.20 However, the copolymers with galactose, fucosylamine, and mannosamine have been targeted to hepatocytes, mouse leukemia L1210 cells, and macrophages, respectively.21 Several specific polysaccharide-based interactions are also known for drug or gene delivery. A sulfated-glucoside polymer activates the fibroblast growth factor,22 which can be used as an active component in tissue-engineering matrixes. Certain modified chitosans are mostly used for hepatocyte and chondrocyte attachment, which works as a carrier material for transplantation or for tissue engineering.23–26 Polysaccharide and protein interaction is facilitated as an enzyme inhibitor27 and in the treatment for infectious diseases.28 Carbohydrate portions of polymers can mimic natural polysaccharides and bind carbohydrate to lectins.29 This high concentrated plant sugar and animal lectins has been used as a matrix for biomolecular purification.30 Ligands of chiral sugar-based polymer binding provide unique matrixes for gel electrophoresis to separate chiral components. Various sugar-based synthetic polymer structures, including linear and branched polymers, comb-like polymers, dendrimers and cross-linked hydrogels, have been reported. Linear polymers are usually linked at the anomeric position of hydroxyl groups of adjacent sugar molecules. Comblike polymers are synthesized from polymerizable sugar derivatives. Dendritic macromolecules, or dendrimers, are synthetic three-dimensional macromolecules prepared from simple branched monomer units. Their unique and monodisperse structure results in improvement of physical and chemical

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Natural-based polymers for biomedical applications

properties when compared to linear polymers. Currently, dendrimers are considered to be one of the nanoscale building blocks for the construction of nano objects, molecular devices, and advanced drug delivery systems.31 Hydrogels are cross-linked polymers which swell significantly in water. Sugar-based hydrogels are hydrophilic and biocompatible and are used in medicine and biomedical engineering such as superabsorbents, contact lenses and matrices for drug delivery systems.32–33 Section 1.2 reviews polysaccharide classes that have medical use.

1.2

Starch

Starch is the major source of energy stored as a carbohydrate in plants. It is composed of two substances: amylose, which is a linear polysaccharide, and amylopectin, which is a branched polysaccharide. Both the forms of starch are polymers of α-D-glucose. Natural starch contains 10–20% amylose and 80–90% amylopectin. Amylose forms a colloidal dispersion in hot water whereas amylopectin is completely insoluble. Starches are hydrolysed to simple sugars using acids or enzymes as catalysts. In hydrolysis of starches, water is used to break long polysaccharide chains into smaller chains or into simple carbohydrates like dextrose. Polydextrose (poly-D-glucose) is a synthetic polymer, formed by heating dextrose with an acid catalyst.

1.2.1

Amylose

Amylose molecules are made up of 200 to 20 000 glucose units, forming a helix structure due to the bond angles between the glucose unit (Fig. 1.1).

1.2.2

Amylopectin

Amylopectin molecules are made up of about two million glucose units. The side chain branches of amylopectin are made up of about 30 glucose units attached with 1α→6 linkages approximately every 20 to 30 glucose units along the chain (Fig. 1.2).

CH2OH O H H H OH H H

OH

CH2OH O H H H O

OH H H

1.1 Amylose

© 2008, Woodhead Publishing Limited

OH

CH2OH CH2OH CH2OH O O H O H H H H H H H H OH H OH H OH H O O O O H

OH

H

OH

H

OH

Polysaccharides as carriers of bioactive agents CH2OH O H H H OH H

O

H

OH H H

CH2OH O H H H OH H O O

OH

CH2OH O H H H

OH

H

CH2OH O H H H O

OH

OH H H

7

OH

CH2OH O H H H

CH2 H H O

O

H

OH

H

H

OH

O

OH

H

H

OH

H O

CH2OH O H H OH

H

H

OH

O

1.2 Amylopectin.

1.2.3

Glycogen

Glucose is stored as glycogen in animal tissues by the process of glycogenesis. Glycogen is a polymer of α-D-glucose similar to amylopectin, but the branches in glycogen are made up of about 13 glucose units The glucose chains are arranged globularly like the branches of a tree originating from a pair of molecules of glycogenin, a protein with a molecular weight of 38 000 that acts as a primer at the core of the structure. Glycogen is easily converted to glucose to provide energy.34–37

1.3

Cellulose

Cellulose is a polymer of β-D-glucose, oriented with –CH2OH groups alternating above and below the plane of the cellulose molecule, thus forming long, unbranched chains (Fig. 1.3). The absence of side chains in cellulose molecules bring them close to each other to form rigid structures. Cellulose is the major structural material of plants. Wood is largely cellulose, and cotton is almost pure cellulose. Cellulose can be hydrolyzed to its constituent glucose units by microorganisms that inhabit the digestive tract of termites and ruminants. Cellulose may be modified in the laboratory by treating it with nitric acid (HNO3) to replace all the hydroxyl groups with nitrate groups (–ONO2) to produce cellulose nitrate (nitrocellulose or guncotton), which is an explosive component of smokeless powder. Partially nitrated cellulose is known as pyroxylin, used in the manufacture of collodion, plastics lacquers and nail polish.34–37

1.3.1

Chitin

Chitin is an unbranched polymer of N-acetyl-D-glucosamine (Fig. 1.4). It is found in fungi and in lower animal exoskeletons, e.g. insect, crab and shrimp

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Natural-based polymers for biomedical applications CH2OH O H H

H O

OH H H H

OH

OH H

OH H H H O CH2OH

OH CH2OH H O H H O H OH H OH H H O H H O H CH2OH OH

O

1.3 Cellulose. CH2OH O H H

H O

NHCOCH3

OH H OH H H H H O H NHCOCH3 CH2OH

H

H O

CH2OH O H OH H H

H O

NHCOCH3

OH H H H H O NHCOCH3 CH2OH

H O

1.4 Chitin.

shells. It is considered as a derivative of cellulose in which the hydroxyl groups of the second carbon of each glucose unit are replaced with acetamido (–NH(C=O)CH3) groups.34–37

1.4

Heparinoid (sulfated polysaccharides)

Heparinoid is a term used for polyionic substances possessing heparin-like effects. Two major polyanionic substances have been studied during the past three decades; one of them is polyanionic polysaccharide. Chemically modified polyanionic polysaccharide includes heparan sulfate, pentosan polysulfate, dextran sulfate or chitin sulfates.38–42 Heparinoid based polysaccharides are long unbranched polysaccharides containing repeating disaccharide units containing either of two amino sugar compounds – N-acetylgalactosamine or N-acetylglucosamine, and a uronic acid such as glucuronate. The heparinoid biological activity is due to the interaction of polysaccharide molecules binding with proteins. Both ionic and hydrogen bonding residues lie in the special manner on the surface of shallow binding pockets on the surface of heparin binding protein.

1.4.1

Heparin

Heparin is a complex mixture of sulfated linear polysaccharide chains present in mast cells (Fig. 1.5). Anticoagulant properties of heparin are depending on the degree of sulfation of the saccharide units. The average molecular weight of heparin is about 12 000 D43–44 consisting of repeating units of trisulphated diasaccharides. It bears an additional number of diasaccharide structures, which makes heparin structure complex.45–48 It is acidic polysaccharide possessing sulfates or N-acetyl groups. The degree of sulfation and the chain

© 2008, Woodhead Publishing Limited

Polysaccharides as carriers of bioactive agents COO H

–

O H

–

CH2OSO3 O H H H OH H O O

H

OH H H

9

–

OSO3

H

–

HNSO3

1.5 Heparin.

size of heparin determine its biological activity. Although heparin has a wide range of biological activities its clinical use is limited in the treatment of blood-clotting disorders as an anticoagulation activity. Heparin may endanger patients due to the high risk of hemorrhage. The discovery of the anti-thrombin III (AT III) pentasaccharide binding site and the elucidation of its structure activity relationship (SAR) demonstrate that heparin possesses a definite sequence within its binding domain which interacts with high specificity and affinity to selected proteins.49 This interaction has been exploited in the development of a highly specific anti-factor Xa agent.50 Heparin may form nonspecific protein interactions with well-defined heparin oligosaccharides. For example, AT III pentasaccharide is used as an anti-factor Xa agent, and interacts with platelet factor 4 (PF-4) and causes some undesired side effects.51–52 The activity of heparin is not due to the heparin molecule but it is due to the sulfate group present in it. Heparin was used in the treatment and prophylaxis of thromboembolic disorders.49,53 Heparin is administered parenterally due to its inability to absorb within the gastrointestinal (GI) tract. Its activity usually occurs within 20–60 minutes after injection with an average half-life period of 1–2 h. The half-life period is reduced in patients with thrombosis disorders and liver impairments. Heparin is bound to the plasma proteins and does not cross the placenta and does not distribute into breast milk. Heparin is excreted in the urine mainly as metabolites although in the administration of large doses, up to 50% may be excreted. Heparin may cause hemorrhage or reversible thrombocytopenia. Heparin drug interaction is established with drugs affecting platelet functions, thrombolytic agents and dihydroergotemine misylete.49,52 Heparin activity in cancer and agiogenesis has been recently studied.54–56

1.4.2

Pentosan sulfate

Pentosan sulfate is an active heparinoid drug in the form of a sulfated chain of xylose sugars linked together (Fig. 1.6). Pentosan is obtained from beechwood shavings and is effectively polyxylose with a molecular weight of approximately 5000, derived from relatively pure lignin derivatives. After sulfation, pentosan, known as pentosan polysulfate (PPS), is a highly sulphated, semi-synthetic polysaccharide somewhat similar to heparin or dextran sulfate.

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Natural-based polymers for biomedical applications –O2C

–O3SO O

O

–O3SHN O

–O2C O

HO

–O3SHN

NHSO3– O

O

O HO

O

O

O –O3SO

OH

OH HO O

–O3SO

–

OSO3

OSO3–

1.6 Pentosan polysulfate.

It is a large and water-soluble molecule, used as a drug since 1960. First, it was used as an anticoagulant (large doses i.v.) then as an anti-inflammatory agent (smaller doses by injection) and for major treatment of interstitial cystitis (oral). Pentosan polysulfate is the polysulfate ester of xylan, a polymer prepared semi-synthetically. The repeating units of the xylan polymer are (1–4) linked β-D-xylopyranoses, with one molecule of the sulfated esters of alpha-Dglucopyranosyluronic acid attached to the 2 position of the xylan approximately after every nine monomeric units.57 The degree and positions of substitution of the sulfate esters and the ring conformation of pentosan polysulfate have been confirmed by 13c-NMR spectroscopy.58 It is very effective in preventing the growth of cancer by stopping the growth of the blood vessels needed for the cancer growth, AIDS infection and in amyloidoses. It is given orally in the form of capsules, which is excreted intact with urine. On continuous administration for few months, it was found on the surface of the bladder and the prostate. About 1.2% patients showed a tendency for increased blood transaminases. Sodium pentosan polysulfate has been used for 40 years for the treatment of a variety of conditions including thrombosis, thrombo-embolic complications, hyperlipidaemia, dyslipoproteinaemia degenerative and diabetic arteriopathies. Its application as a DMAOD has attracted attention recently. Recently other derivatives such as calcium pentosan polysulfate (CaPPS) have been investigated and found to exhibit higher oral bioavailability than sodium salt.59 The PPS medicinal applications include anticoagulant, fibrolytic and anti-inflammatory agents.60–62 It may be used as hypolipidemic agent,38,41,63 in reduction of smooth muscle cells proliferation, and as inhibitor of enzymatic activity including heparase, protein kinase, and reverse trascriptase.41 It is used as angiostatic and potent anti–HIV agent (in vitro).64 PPS was found effective as an anti-prion agent.65

1.5

Dextran

Dextran is a polysaccharide macromolecule (Fig. 1.7) used for selective transport and is a carrier for a wide range of therapeutic agents due to its

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Polysaccharides as carriers of bioactive agents

11

CH2 O H H H OH H HO O H HO CH2

CH2 O H H H OH H HO O 6

H HO CH2 O H H 5 H O H H H OH H 4 1 HO O H HO O H HO 2 3 H HO CH2 O H H H OH H HO O H HO

1.7 Dextran.

excellent physico-chemical properties and physiological acceptance. Dextrans have been used for drug targeting, increasing blood circulation time, stabilization of therapeutic agents, solubilization of drugs, reduction of side effects, sustained release action and depot properties.66 Dextrans of different chemical composition are synthesized by a large number of bacteria of family Lactobacillaceae and mainly from Leuconostoc mesenteroids, Leuconostoc dextranicum and Streptobacterium dextranicum. The synthesis of low molecular weight dextran is done in sucrose or other carbohydrate mediums containing anhydro-D-glucopyranose units.67 Dextrans derived from Leuconostoc mesenteroids NRRL B-152 are of particular pharmaceutical interest. The microbiological synthesis product is known as ‘native-dextran’. Clinical dextrans are obtained from high molecular weight native dextrans after their partial depolymerization by acid hydrolysis and fractionation.68 Dextrans obtained from different sources possess different structures and properties, i.e. degree of branching, relative quantity of particular type of glycosidic links, molecular weight, solubility, optical activity and physiological action. Dextrans are soluble in water, formamide and dimethylsulfoxide and insoluble in alcohol and acetone. Native dextran is a polymer of high molecular weight ranging between 107 to 108. Its molecular weight is reduced by acid hydrolysis irrespective of the nature of acid used. Native dextran also possesses a high degree of polydispersibility. The optical rotation of different aqueous solutions of dextrans varies from +199° to +235°. The viscosity is affected due to the degree of branching, the nature and pH of solvent, the number of intermolecular bonds and temperature.69 Dextran is the name of a large class of a-D-glucans with anhydro-Dglucopyranose units. a-1,6-linkages are predominant features of dextrans.70 Dextrans are composed of 95% a-1,6-glucopyranosidic linkages and 5% 1,3-

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Natural-based polymers for biomedical applications

linkages. The 1,3-linkages are the points for the attachment of side chains, of which about 85% are 1 or 2 glucose residues in length and the remaining 15% of side chain may have an average length of 33 glucose residues. Dextrans contain different α-1,2-, α-1,3- and α-1,4-glycosidic bonds, by which means the side chains are usually attached to the main chains. Dextrans are extensively used as backbones for attaching drugs. Dextrans form alkoxide dextranates on reaction with alkali and alkaline earth metals.71 Oxidation products of dextrans are useful in preparation of several new derivatives of dextrans. A water-soluble chlorodeoxydextran has been prepared by treating dextran with thionyl chloride in DMF. Aminodeoxy derivatives have also been obtained by nucleophilic substitution reactions.72 A deoxymercapto derivative is obtained by pyrolysis of dextran-xanthate with sodium nitrite followed by treatment of the resulting polymer with alkali.73

1.5.1

Pharmacokinetic fate of dextran

Various physico-chemical properties of dextran, such as molecular size and shape, flexibility, charge, hydrophilic lipophilic balance, are determinants for the pharmacokinetic fate of dextran.74 When given parenterally, intravascular persistence of dextran varies dramatically with molecular weight. Dextrans of molecular weight less than 70 000 have rapid elimination rates during the first hour after injection followed by a slower decrease in concentration.75 Dextrans with molecular weights in the range 50 000–70 000 show prolonged survival in the circulatory system. The high polarity of dextrans excludes their transcellular passage and their size prevents the passage through the gastrointestinal tract. Dextrans are depolymerised by the enzyme dextranase present in the intestine.

1.6

Pectin

Pectin is a polysaccharide that acts as a cementing material in the cell walls of all plant tissues. The white portion of the rind of lemons and oranges contains approximately 30% pectin. Pectin is the methylated ester of polygalacturonic acid, which consists of chains of 300 to 1000 galacturonic acid units joined with 1α→4 linkages (Fig. 1.8). The degree of esterification affects the gelling properties of pectin. The structure shows three methyl COOCH3 O

O H

H H O OH H H

H

OH

COOH

COOCH3

O H H O OH H

O H H O OH H

H

1.8 Pectin.

© 2008, Woodhead Publishing Limited

OH

H

H

OH

COOH

COOCH3 O H H H

OH H H

OH

O H

O H H O OH H H

OH

Polysaccharides as carriers of bioactive agents

13

ester forms (–COOCH3) for every two carboxyl groups (–COOH), hence it is has a 60% degree of esterification. The substituted residues at C-4 with neutral and acidic oligosaccharide side chain are composed of arbinose, galactose, fructose and glucuronic acid.76 Pectin increases viscosity and volume of stools and hence is used against constipation and diarrhea. Pectin is also used in throat lozenges. It is also used in wound healing preparations and in several special medical adhesives, such as colostomy devices. In cosmetic products pectin works as a stabilizer. In ruminant nutrition, depending on the extent of lignification of the cell wall, pectin is up to 90% digestible by bacterial enzymes. Ruminant nutritionists recommend that the digestibility and energy concentration in forages is improved by increasing pectin concentration in the forage. The gelling, binding, biocompability and nontoxicity properties of pectin make it a promising biopolymer to construct drug carriers for controlled drug delivery. Various drugs can be incorporated into pectin formulations with high loading efficiency using simple procedures. Various chemical compositions of pectin are used for several specific applications. Highly polar pectin derivatives can penetrate deeply into tissue to prolong the residual time to incorporate drugs and enhance their penetration. Pectins and zein composite gels are able to deliver a drug to specific GI segments at the desired time.77

1.7

Arabinogalactan

Arabinogalactans (AG) are a class of long, densely branched polysaccharides with molecular weights ranging from 10–20 kDa. In nature, arabinogalactans are found in a wide range of plants; however, the primary source of AG is the larch tree, and the most available arabinogalactan is from the western larch (Larix occidentalis), which provides a rich harvest of free arabinogalactan from its inner bark. O O

OH

OH O

OH

OH

O

OH O

OH OH

O

O O OH

OH

O OH

OH O

OH

O OH

1.9 Arabinogalactan.

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Natural-based polymers for biomedical applications

AG from the western larch is a branched, water-soluble polysaccharide with a relatively narrow molecular weight distribution. The basic building units of AG are arabinose and galactose in a ratio of approximately 1:6 (Fig. 1.9). The backbone contains D-galactopyranose residues linked by β(1–3) bonds. The majority of these residues bear branches consisting of either one (23%) or two (46%) D-galactopyranosyl residues linked by β(1–6) bonds. Smaller percentages of main chain residues bear larger branches with terminal arabinose residues.78 High solubility in water (70%), biocompatibility, biodegragrability and ease of conjugation in aqueous medium makes AG an attractive polymer for biomedical applications.

1.7.1

Arabinogalactan toxicity

Larch arabinogalactan is a safe and effective immune-stimulating phytochemical. It is used as a dietary fiber and in food applications. The acute toxicity of arabinogalactan was investigated by Groman.79 AG did not cause mortality in either rats or mice injected intravenously with 5 g/kg AG, nor were there signs or symptoms of toxicity evident in either species during the in vivo phase of the study. Repeat dose toxicity was evaluated after the injection of 31–500 mg/kg/day of AG to rats. There were no overt clinical signs or symptoms of toxicity related to AG administration and the animals gained weight over the 90 day dosing period. AG binds to the asialoglycoprotein receptor in its naturally occurring form and therefore can be useful in the hepatic delivery of diagnostic or therapeutic agents.

1.7.2

Arabinogalactan pharmacokinetics

The clearance of AG determined by injection of [3H] AG to rats, showed an elimination of AG from the blood with half-life of 3.8 min at 30 min postinjection; 31% of the injected dose was found in the liver and the radioactivity remaining in the liver declined with a half life of 3.4 days. The reason for the strong interaction of AG with the asialoglycoprotein receptor may lie in the highly branched structure and numerous terminal galactose or arabinose residues.

1.7.3

Arabinogalactan clinical indications

AG used in various clinical studies to provide different medical actions. Larch arabinogalactan is an excellent source of dietary fiber that is able to increase short-chain fatty acid production (primarily butyrate) via vigorous fermentation by intestinal microflora.80 In addition, it increase levels of beneficial intestinal anaerobes, particularly Bifidobacterium longum, via their fermentation specificity for arabinogalactan compared to other complex carbohydrates.81,82 In cancer therapies, larch arabinogalactan may be used as

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Polysaccharides as carriers of bioactive agents

15

an effective adjunct due to its ability to stimulate NK cell cytotoxicity, stimulate the immune system, and block metastasis of tumor cells to the liver.80 Stimulation of NK cell activity by larch arabinogalactan and decrease in NK cell cytotoxicity has been associated with recovery in certain cases of chronic fatigue syndrome,83 viral hepatitis (hepatitis B and C)84 and in the case of multiple sclerosis.85

1.8

Drug conjugated polysaccharides

Targeted drug delivery, based on macromolecular polysaccharides, attracts significant attention due to their ability to improve the pharmacokinetics and pharmacodynamics for small drug, protein and enzyme molecules. The attachment of therapeutic agents to polysaccharide molecules by conjugation leads to an increase in the duration of activity. There are two major types of polysaccharide-drug conjugate. Conjugates of small molecule drugs to polysaccharides rendering the conjugated drug inactive are called macromolecular prodrugs, which need to release the active drug in vivo in order for the drug to exert its pharmacological actions. On the other hand, conjugation of large molecular weight therapeutic agents such as peptides and proteins with polysaccharides usually results in conjugates, which retain partial or complete activity. Physicochemical properties of polysaccharides such as molecular weight, structure, and charge, significantly impact the pharmacokinetic/dynamic properties of the macromolecule–drug conjugates. The chemical structures of various polysaccharides are discussed in this chapter. All these macromolecules are basically neutral in nature. However, chemical modifications result in positively- or negatively-charged macromolecules. In such conjugations the drug molecule is covalently attached to the polysaccharide macromolecule directly or with a spacer arm or linker known as a prodrug. This macromolecular prodrug is normally inactive and is expected to be relatively stable in vitro and releases the active drug at the specific site in vivo. These prodrugs may be used for their systemic or local effects.

1.8.1

Systemic effects

The most widely investigated applications of macromolecular polysaccharide conjugates are in the area of cancer chemotherapy (Table 1.1). This is mostly accomplished via passive targeting of the macromolecular prodrug to the tumor. The importance of passive targeting is usually long half lives of the prodrugs in circulation, as opposed to shorter residence times of the drugs themselves, accompanied by an enhanced permeation and retention (EPR) effect by the tumor. The latter is due to the increased permeability of the tumor vasculature, resulting in increased prodrug entry into the tumor tissue,

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Natural-based polymers for biomedical applications

complemented by a decreased lymphatic drainage of the tumor, which causes the retention of the prodrug. Hence a prodrug, having a longer plasma halflife than a drug, gradually accumulates in the tumor site, resulting in passive targeting. On the other hand, for active targeting, tumor or tissue specific ligands are attached to the macromolecular prodrug, which is actively taken up, by the tumor and/or the desired tissue. The release of the active drug at the specific site is essential for effective prodrugs. Consequently, the nature of the linkage between the macromolecule and the drug becomes crucial.86 Earlier workers used linkers for conjugations which are susceptible to chemical hydrolysis for release of the drug. For example, α-aminobutyric acid, αaminocaproic acid, and α-aminocaprylic acid with 4, 6, and 8 carbon atoms, respectively, are used as linkers between mitomycin C and dextran.87 Results found that the released half life of mitomycin at the pH 7.4 buffer (37°C) increased with increase in the length of the linker.87 It happened because the drug release rates from the prodrug in plasma and liver homogenates were not significantly different than those in the presence of plasma, suggesting the release of mitomycin C is based on chemical hydrolysis process.88 More recently, linkers are being designed to release the drug specifically in the lysosomal compartment, where the macromolecular prodrug disintegrates completely after endocytosis. It is due to lysosomes containing various types of enzymes such as glucosidases, esterases, and proteases, which makes release of drugs easier from macromolecular prodrugs.

1.8.2

Arabinogalactan-amphotericin B conjugates

The arabinogalactan-amphotericin B conjugates (Fig. 1.10) have been synthesized as a new drug moiety providing higher water solubility and lower toxicity.89 Amphotericin B (AmB), a polyene antibiotic, is a standard drug for the treatment of fungal infections90 and is currently recommended as a second-line treatment for visceral leishmaniasis and mucocutaneous leishmaniasis,91 especially with human HIV coinfection. However, AmB therapy is limited due to its negligible solubility in aqueous solution and poor solubility in most organic solvents, and due to its toxicity, mainly to the kidneys, central nervous system, and liver, and side effects such as nausea, fever, and chills.92 AG was oxidized using potassium periodate, purified from the oxidizing agent using ion-exchange chromatography, and reacted with AmB to form the Schiff base. The Schiff base and aldehydes were reduced to the amine and hydroxyl respectively using borohydride. All reactions took place in aqueous media. Both amine and imine AmB–AG conjugates were soluble in water and exhibited improved stability in aqueous solutions as compared to the unbound drug. The conjugates showed comparable minimum inhibitory concentration (MIC) values against Candida albicans. The conjugates were

© 2008, Woodhead Publishing Limited

Polysaccharides as carriers of bioactive agents Table 1.1 Polysaccharides as carriers in macromolecular prodrugs of anticancer agents74 Polymer

Mol. Wt.

Drug

Targeting method

Comments

Dextran

40 kD

Doxorubicin (DOX)

Passive

DOX was linked to Dextran through Gly-Leu-Gly tripeptide or via a hexamethylene spacer. In the presence of papain, the tripeptide conjugate released 43% of its DOX content in 48 h. However, the conjugate with hexamethylene spacer did not release any DOX. In vitro studies showed that while the conjugate with the tripeptide linker was more effective than the one with the hexamethylene spacer in Hep-3B hepatoma cells, both were ineffective against SiHa cells which lack lysosomal enzymes.

Carboxymethylpullulan

150 kD

Doxorubicin (DOX)

Passive

DOX was connected to CMP using Gly-Gly-Phe-Gly (1), Gly-Phe-Gly-Gly (2), or GlyGly-Gly-Gly (3) linkers or by direct linking (4). All conjugates were more stable than free drug at pH 7.4. The in vivo release of DOX from conjugates and their antitumor activity were dependent on the type of linker for conjugates 1-3. Conjugate 4 (no linker) did not release DOX, nor was it effective in vivo.

Dicarboxymethyl dextran

42 kD

Cisplatin

Active: Via incorporation of branched galactose units

In vitro studies using HEPG2 human hepatoma cells showed that the active targeting of the conjugate using branched galactose significantly increased the effectiveness of cisplatin, compared with those of passive targeting.

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18

O O OH

OH

O O

OH

O

O

O

KIO4

O

O

O OH

OH

OH

O OH

OH

O

O

OH

Dowex-1

O

OH

O

OH

OH

O O

O

OH

OH

O

O OH

O O

OH

O

Arabinogalactan

O

O

O

O

Oxidized arabinogalactan O OH

OH

O

O

O O

OH O

OH

O

OH

OH

OH

O

OH

O

HO

O O

O

OH

OH

OH

OH

OH

O

NH2

OH O

O

O OH

O O

O O

OH

O

O O

OH OH O

OH

OH

OH

OH

N

OH COOH

O

O

O

O

O

O

OH

O

OH O

OH

OH

OH

OH

COOH

OH

OH O

1.10 Synthesis of arabinogalactan-amphotericin B conjugates. © 2008, Woodhead Publishing Limited

NH

OH HO

O

OH

OH

OH

20 h

AmB-AG imine conjugate

OH

O

OH

NaBH4

OH

O

O

OH

O

OH OH

O O

O

OH

OH

OH O

OH

O

O

OH

O

O

OH

OH

OH O OH

HO

OH

Amphotericin B

O

O

O

O

O

O

O

O

B.Borate 0.1M, pH-11 48 h

COOH

OH

AmB-AG amine conjugate

O

OH

Natural-based polymers for biomedical applications

OH OH

O

OH

OH

Polysaccharides as carriers of bioactive agents

19

about 60 times less hemolytic against sheep erythrocytes than the free drug, and about 40 times less toxic in BALB/c mice.

1.8.3

Local effects

Polysaccharides are used for local delivery of anti-inflammatory agents to the colon diseases such as colitis and Crohn’s disease. The synthesis93–94 and in vitro94–98 and in vivo98–100 release of a dextran-nonsteroidal antiinflammatory drug (NSAID) ester conjugate formed by direct conjugation of dextrans with NSAIDs is an example of local delivery of polysaccharide conjugates with an antiinflamatory drug. These studies demonstrated that after oral administration, the enzymatic release of NSAIDs from dextran-NSAID conjugates would occur in cecum and colon. However, the release of NSAIDs in the upper part of the gastrointestinal tract is slow and occurs by chemical hydrolysis. It is because of the large molecule of dextrans, esterases in gastrointestinal tract could not hydrolyze the conjugates.96–97 However, enzyme dextranases in the colon reduce the molecular weight of dextrans, making them more susceptible to the esterase action. This study served as a model for the use of dextrans in colonic delivery of other drugs like corticosteroids,101–103 after the oral administration of dextran–drug conjugates.

1.8.4

In vivo disposition of carriers

The macromolecular carrier itself mainly dictates the pharmacokinetics of polysaccharide macromolecular–drug conjugates. Thus an understanding of the disposition of polysaccharides is crucial in the designing of proper polysaccharide-based delivery systems. Some of the polysaccharide drug conjugates are now discussed.

1.9

Polysaccharide dextrans

Polysaccharide dextrans are the most widely studied drug conjugates in terms of their in vivo disposition with regard to molecular weight, charge and dose. A brief overview of the disposition of dextrans is discussed here.

1.9.1

Native dextran

High molecular weight dextrans are not substantially absorbed on oral administration.104 Thus dextran–drug conjugates designed for systemic effects need to be administered by injection routes (sc, im, or iv). However, dextran prodrugs may be used orally for their local effects in the gastrointestinal tract as discussed earlier. The systematic determination of pharmacokinetics and the tissue distribution of fluorescein-labeled dextrans (FDs) of different

© 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications

molecular weights [4 kD (FD-4), 20 kD (FD-20), 70 kD (FD-70) and 150 kD (FD-150)] has been performed in rats.105–107 Though the dextrans do not have a chromophore in their structure, the fluorescein label was used for sensitive detection of dextrans by a fluorescence detector using a high performance, size exclusion chromatographic method. These studies revealed that after iv administration of a single 5-mg/kg dose of FDs, the serum concentrations of FD-4 and FD-20 declined rapidly, and of FD-70 and FD-150 persisted much longer. The molecular weight dependency of the serum concentrations of FDs were attributed to a molecular weight dependent renal clearance of the macromolecule; whereas the renal clearances of FD-4 and FD-20 were relatively high, and of FD-70 and FD-150 were found to be negligible.105 This kinetic behavior is reliable with the glomerular capillary walls pore sizes excluding dextrans with a radius of 4.4 nm (MW > 40 kD) allowing unlimited excretion of dextrans with a radius of 2 nm (MW, < 10 kD).108 In contrast to changes in renal clearance increased molecular weight generally resulted in a greater accumulation of FDs in the liver and spleen.105 Except for FD-4, the amounts of FDs found in the liver were very high, even at extended times after the administration of the macromolecule. This means an increase in the molecular weight of FDs from 4 kD to 20 kD resulted in an increase in the accumulation of dextran in tissue. The same trend was observed when the molecular weight of FD was increased to 70 kD. However, when the molecular weight was further increased from 70 kD to 150 kD, no increase in the amount of FD in the liver was observed. Similar molecular weight dependency was also observed for the spleen. Additionally, the concentrations of FDs in the spleen were found to be relatively high because of the small weight of the spleen, compared to the liver. The percentage of the dose found in this organ was significantly less than that found in the liver. For the liver, the ratio increased by > 40-fold when the molecular weight increased from 4 kD to 20 kD. Further increase in the molecular weight from 20 kD to 70 kD resulted in a modest increase (<2 fold) in the ratio. However, when the molecular weight was increased from 70 kD to 150 kD, the liver:plasma ratio declined by > 3 fold. The high concentration of FD-70 in the liver resulted in recovery of ~60% of the administered dose (5 mg/kg) in this organ.105 For the spleen, an increase in molecular weight from 4 kD to 150 kD resulted in a progressive increase in the tissue:plasma ratio. Aside from the liver and the spleen, the concentrations of FDs were low in the other studied tissues except kidneys, which showed high concentrations of low molecular weight dextrans. The concentrations of FDs in the brain were zero, which could be very important if dextrans are to be used for delivery of drugs for which brain is the site of toxic, rather than desirable effects (such as immunosuppressants109–110). Based on these data, FD-70 was highly targeted to the liver and spleen. Additional studies94 investigating the dose-dependency of FD-4 and FD-150 have shown that the kinetics of renally excreted FD-4 are linear, whereas modest non-linearity is

© 2008, Woodhead Publishing Limited

Polysaccharides as carriers of bioactive agents

21

observed in the hepatic accumulation of FD-150; when the dose was increased 100 fold from 1 to 100 mg/kg, the percentage dose recovered in the liver decreased from 68.5% to 41.5%.106 Further, the plasma clearance of FD-150 decreased by a factor of 2 when the dose was increased from 1 to 100 mg/kg. The data suggest the elimination of dextrans is dependent on the hepatic accumulation, rather than renal excretion. Nonlinearity in the kinetics is only expected at higher doses of dextran.

1.9.2

Chemically-modified dextran

Native dextrans were modified to introduce electrical charge for passive delivery and/or to attach ligands for receptor-mediated drug delivery. Introduction of negative charges to dextrans by carboxymethylation decreases the macromolecule uptake by the tissues and increases their time in circulation.111–113 On the other hand, positively charged dextrans (e.g. DEAE dextrans) are rapidly cleared from the circulation and taken up by tissues, notably by the liver.111–113 During the studies on the conjugates of an analog of the anticancer drug camptothecin with carboxymethyldextran (CMD), the effects of degree of carboxymethyl substitution on the in vivo pharmacokinetics of the carrier itself were investigated.114 After iv injection of single doses of 20 mg/kg to rats bearing Yoshida carcinoma, the plasma AUCs of the 110 kD CMDs with degrees of substitution (DS) of 0.2, 0.4 and 0.6 were similar. However, an increase in the DS to 1.0 resulted in an AUC value half that of CMD with a DS of 0.6. Nevertheless, the AUCs of all the studied 110 kD CMDs were three- to six-fold larger than that of neutral dextran with a molecular weight of 150 kD, due to lower clearance of CMDs. In terms of distribution into liver, accumulations of 110 kD CMD with a DS of 0.2–0.6 were substantially lower than that of neutral dextran with a MW of 150 kD.114 However, a further increase of DS to 1.0 resulted in a sharp increase in the liver accumulation to a level comparable to the neutral dextran 150 kD. These studies indicate that there is no linear relationship between the degree of carboxymethylation and hepatic accumulation of CMDs in tumorbearing rats. However, the extents of tumor accumulation of CMDs with DS of 0.2–1.0 were comparable to each other and higher than that of neutral dextran 150 kD.114 The effect of chemical modifications on the kinetics of CMD molecular weights. Oxidation of CMD by the sodium periodate method, used for conjugation of dextrans with amine containing drugs, resulted in > 15-fold reduction in the AUC of the macromolecule in rat plasma.115 Conjugation of the oxidized CMD with the relatively hydrophilic ethanolamine decreased the AUC of the macromolecule by an additional five-fold. On the other hand, attachment of the hydrophobic analgesic DA5018 to the oxidized CMD resulted in a seven-fold increase in the AUC of the oxidized CMD. The investigators show the chemical modification of CMD resulted in a

© 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications

reduction in the effective molecular size of the macromolecule, corresponding to the decreases in the AUC of the modified CMDs.115 The asialoglycoprotein receptors of hepatocytes and mannose receptors of non-parenchymal hepatic cells have been exploited for cell-selective delivery of therapeutic agents using galactosylated or mannosylated CMD prodrugs. Studies in rats116 and mice117 show the after iv injection, galactosylated and mannosylated CMD accumulated predominantly in liver parenchymal and nonparenchymal cells, respectively. This approach has been used for targeted delivery of Ara-C and cisplatin.118

1.10

Mannan

In vivo disposition of mannan and its derivatives is dependent on tissue mannose receptors and serum mannan binding proteins. Mannose receptors show good affinity toward the mannose polymers like mannan and are abundant in liver endothelial and Kupffer cells and in spleen and alveolar macrophages.119–120 In serum, mannan binds to mannan binding protein (MBP), which is present in both humans121 and animals.122–125 The disposition of Candida albicans mannan (CAM) and Cryptococcus neoformans glucuronoxylomannan (GluXM) has been achieved by intravenously injecting doses of 20 mg and 20 µg, respectively, to rabbits.126 The two mannans are different in terms of their side chains, with CAM containing α-(1-2)- and α(1-3)-linked mannose units and GluXM containing α-(1-2)-linked glucuronic acid and α-(1-2)-linked xylose. For CAM, both free and protein-bound macromolecules were detected in serum. However, only protein-bound GluXM was detected in the serum of rabbits. Whereas CAM (free and bound) showed a short serum half life of 2 h, the half life of bound GluXM from Cryptococcus neoformans was much longer (close to 24 h).126 Additionally, only CAM was excreted in the urine of rabbits. The faster serum disappearance of CAM was attributed to the presence of mannose units on its surface, as opposed to the presence of glucuronic acid and xylose, which do not bind to mannose receptors, on the surface of GluXM. The investigation on the tissue accumulation of CAM in mice after intravenous injection of 200 µg of the macromolecule shows the accumulation of doses in the liver and spleen of mice, where the carbohydrate persisted for 97 days.126 In a subsequent study, the relative distribution of CAM in mice at 90 min after the injection of a 5 mg dose was blood > liver > lung > spleen was reported.127 However, after this relatively low dose, most of CAM in blood was bound to mannose binding proteins; only ~1% of the dose was free in blood. Thus it is concluded that the significant accumulation of the macromolecule in liver, lung, and spleen should be due to an active binding process.127 This is in agreement with the presence of mannose receptors in these tissues. However, recent data suggest that MBP in serum may act against the receptor-mediated tissue

© 2008, Woodhead Publishing Limited

Polysaccharides as carriers of bioactive agents

23

accumulation of mannose containing macromolecules by trapping them in blood. The presence of MBP in cultured mouse peritoneal macrophages dose-dependently reduced their transfection by DNA/mannosylated liposome complexes.128 Additionally, studies in mice show an increase in the dose of ligands leading to saturation of the mannose receptors, causing a decrease in the percentage of the dose accumulated in the liver.129 Overall, the trapping effects of serum MBP at low doses and saturation of the tissue (especially the liver) mannose receptors at higher doses cast doubt on the clinical utility of this approach for targeting drugs and macromolecules to the liver.

1.11

Pullulan

1.11.1 Native pullulan 125I-labeled pullulan and dextran of different molecular weights were used to investigate the molecular weight-dependency of the disposition of polysaccharides in mice.130 The results found, similar to dextrans, an increase in the molecular weight of pullulan from 6 kD to 190 kD was associated with a progressive increase in the plasma AUC and liver accumulation of pullulans. Recently, using pullulan with a molecular weight of 60 kD reported a nonlinear disposition of pullulan in rats.131 An increase in dose from 1.5 mg/kg to 24 mg/kg caused an increase in plasma concentration of the macromolecule, which was responsible for a drastic decline in its hepatic accumulation.

1.11.2 Chemically-modified pullulan Plasma and tissue disposition of a negatively charged pullulan, carboxymethylpullulan (CMP), after intravenous administration of single 10-mg/kg doses of the radiolabeled macromolecule to tumor-bearing rats was studied.132 Like dextran and native pullulans, an increase in the molecular weight of CMP from 24 kDa to 100 kDa resulted in a progressive increase in the persistence of the macromolecule in plasma. However, CMP showed more persistence in the circulation and relatively lower accumulation in the liver,132 when compared with the native pullulan.130–131 A pharmacokinetic comparison of radiolabeled pullulan and CMP (MW: 150 kD) in rats after the administration of a single 1-mg/kg dose has been done.133 The plasma AUC of pullulan found increased by > 30-fold when it was carboxymethylated. The increase in the plasma AUC was responsible for > 100-fold decrease in the liver uptake clearance of CMP. Although carboxymethylation also decreased the uptake clearance of the macromolecule into the spleen, the decrease was only ~3-fold. Consequently, carboxymethylation changed the selectivity of the macromolecule from liver (pullulan) to spleen and blood (CMP).

© 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications

Furthermore, at 24 h after the injection, accumulation of CMP in lymph nodes was found substantially greater than the native macromolecule.133 The small dose (1 mg/kg) used precludes extrapolation of the data to CMP prodrugs for which higher doses of the carrier are needed.132,134 Additionally, whether CMP is devoid of nonlinear pharmacokinetics, observed with pullulan, remains to be fully investigated. Several polysaccharides have been used in different ways in the field of drug delivery.

1.12

Polysaccharide macromolecule–protein conjugates

Some proteins or enzymes retain their activity partially, by their covalent attachment to polysaccharides. The polysaccharide conjugations may be used to prolong the in vivo residence time and the effects of these proteins. For example, dextran conjugates of the anticancer enzymes asparaginase135 and carboxypeptidase G2136–137 achieve significantly longer plasma half lives, resulting in prolongation of enzymatic activity. In unconjugated form, the immunogenicity of the enzymes/proteins or xenogenic antibodies normally results in their rapid removal from the body and the possibility of allergic reactions after multiple doses. Covalent binding of these proteins to dextran could potentially alleviate both of these problems. For example, trichosanthin (TCS), a protein that induces abortion and inhibits the growth of choriocarcinoma and replication of HIV-1 is significantly antigenic. However, attachment of dextran to a potential antigenic site of TCS significantly reduces both IgE and IgG response to the protein, retaining 50% of its abortifacient activity.138 Similarly, vaccines conjugated to pullulan polysaccharides show reduced IgE response while retaining the neutralizing antibody production property of the unconjugated vaccine, by reducing the side effects.139–140 It has been shown that some gliomas, melanomas, and squamous carcinomas over-express the epidermal growth factor (EGF) receptor. Therefore, radio labeled EGF may be used for radiotherapy of these tumors. However, the residence time of EGF in the tumor cells is short. Conjugated EGF with dextran 20 kD showed binding to EGF receptor.141 The 125I labeled conjugate showed more than 20 h of residence in human malignant glioma cells, while the free EGF was removed very rapidly.141 Overall, these studies indicate the binding of EGF–dextran conjugates to the EGF receptors is specific because free EGF inhibits this type of binding. Additionally, the effects of 125I-EGFDextran, EGF-Dextran-125I, and 125I-EGF-Dextran-125I on the attachment of the toxic radio nuclides to the dextran carrier may result in more radioactivity exposure than when the radio nuclide is attached to EGF.142 Clinical studies are currently underway to take advantage of this conjugate for cancers associated with a significant over-expression of EGF receptors.142

© 2008, Woodhead Publishing Limited

Polysaccharides as carriers of bioactive agents

1.13

25

Cationic polysaccharides for gene delivery

Currently gene therapy is widely investigated for use in cancer, AIDS, and cardiovascular diseases.143 However, the clinical applications of gene therapy require the development of safe and efficient delivery vectors in vivo. Gene delivery for therapeutic applications involves two strategies: corrective or cytotoxic gene therapy. The first approach includes correction of genetic defects in target cells. This strategy is exploited for the treatment of diseases with single gene disorders.144 The second approach includes destruction of target cells using a cytotoxic pathway. This strategy is used for treatment of uterine leiomyomata and of malignant tumors, including ovarian, breast and endometrial carcinoma.144 Gene’s administration for therapeutic purposes can be done using several techniques including direct introduction of transgenes by cell electroporation, microinjection of DNA, and incorporation of the gene by viral or nonviral vectors in vivo or ex vivo. In vivo delivery of the transgene is done directly by administration of the gene or by using a vector as gene carrier into the patient or at the target organ, and, effectively applied to any cell. Ex vivo administration includes harvesting and cultivation of cells from patients with in vitro gene transfer and reintroduction of transfected cells. The potential target cells for the transfection include lymphocytes, bone marrow cells, umbilical cord blood stem cells, hepatocytes, tumor cells, and skin fibroblasts.147 The main goal of gene therapy is to deliver DNA to target cells accompanied by a high level of desired gene expression. DNA can be delivered into the cell nucleus directly by injection or via specific carriers. Gene carriers are divided into three main groups: viral carriers,145 in which delivered DNA is inserted into the viral genome; physical means; and synthetic vectors.146 The success of gene therapy is attributed to the efficiency of the delivery system used to transport the materials into the nucleus and the stability of the achieved transfection.

1.13.1 Gene delivery systems Gene delivery to the cell proceeds through the following general pathway: formation of the DNA-containing particles, uptake of the particles into the cell, entrance of the particles into the cytoplasm, transport of intact DNA to the nucleus, and finally, expression of the delivered gene. The delivery process may fail at any one of these steps, resulting in reduced transfection efficiency. Viruses have evolved mechanisms that proceed via these steps, easily allowing the DNA to reach the nucleus at high yields. Although viral vectors allow a high transfection rate of the foreign material inserted in the viral genome, their major drawback is the anti-vector immunity, which restricts the administration of repeated doses. In addition, limited capacity to carry DNA,

© 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications

short shelf life, toxicity, inflammatory responses, insertional mutagenesis, and oncogenic effects can occur in vivo147 restricting viral use in gene therapy. The limitations of viral vectors have led to the evaluation and development of alternative vectors based on nonviral systems.

1.13.2 Cationic nonviral vectors Appropriate molecular weight polymers are specifically designed for coupling cell or tissue-specific targeting moieties. Cationic carriers are accepted widely due to their ability to condense DNA and interact with the cell efficiently.148 Cationic polymers used for nucleic acid delivery acquire their charge from primary, secondary, tertiary and quaternary amino groups. Such polycations exhibit a random distribution of cationic sites along the polymer chain. Polyplexes (DNA/polycation complexes) form spontaneously due to the electrostatic interaction between anionic phosphate groups of the DNA and positively charged groups of the polycations. The mechanism of gene transfer across the cell membrane is not clearly understood. Cationic polymers are relatively poor in carrying DNA molecules across the membrane compared to viral vectors. To reach cells, the complexes must diffuse through the capillary network, escape macrophages, and interact with the cell membrane.149 They must be internalized through endocytosis and then exit the endosome in the cytoplasm, reach the nucleus and be transcripted.150 Cationic polymer systems have several advantages over virus vectors, e.g. low immunogenicity and easy manufacture. They form complexes with DNA and protect it against nuclease degradation.148 Cationic polymers are used to condense and deliver DNA both in vitro and in vivo. Among the large number of polycations used in gene delivery, cationic polysaccharides (including chitosan and their derivatives) and polysaccharide-based oligoamine derivatives has been discussed. Most of the polycations are toxic to cells and are nonbiodegradable.151 Cationic polysaccharides are considered the most attractive candidates for gene delivery. They are relatively nontoxic, biodegradable, and biocompatible materials simply modified for improved physicochemical properties.152–153

1.14

Diethylaminoethyl-dextran

Diethylaminoethyl-Dextran (DEAE-Dextran) is a polycationic derivative of dextran prepared by reacting diethylaminoethyl chloride with dextran in basic aqueous medium.154 DEAE-Dextran is formed from two types of subunits: the single tertiary DEAE-group and tandem groups with a quaternary amine group. The quaternary group is strongly basic (pKb 14), whereas the tandem DEAE-group has a pKb of 5.7 and the single DEAE-group has a pKb of 9.5. DEAE-Dextran was one of the first chemicals used for the delivery of transgenes

© 2008, Woodhead Publishing Limited

Polysaccharides as carriers of bioactive agents

27

into cultured mammalian cells.155 Positively charged DEAE-Dextran can associate with negatively charged nucleic acids to form a DNA/polymer complex.156 The cationic nature of the polyplexes, interacts with a negatively charged cell membrane.157 Uptake of the complex presumably takes place by the endocytosis process. All plasmids enter into the cell using DEAEDextran mediated gene transfer assembled into nucleosome-containing minichromosomes. The DEAE-Dextran polymer is a suitable candidate to deliver nucleic acids into cells for transient expression.155,158,159 Several studies found the use of DEAE-Dextran for DNA transfection by colon epithelial cells in vivo. Chloramphenicol acetyltransferase (CAT) activity is examined in the transfected colon segments using DEAE-Dextran, liposomes, and calcium phosphate transfection systems.160 Expression levels of CAT after in vivo DEAE-Dextran mediated gene transfer increased with higher doses of transfected DNA. The transfection is surprisingly achieved by DEAEDextran, which was at least as effective as liposome-mediated gene transfer. DEAE-Dextran allowed superior transfection in the transfer of DNA to human macrophages,155 its transfecting efficiency in a wide range of cell lines is still very low in comparison to other cationic vectors such as Polybrene®, PEI, dendrimers, etc.

1.15

Polysaccharide–oligoamine based conjugates

A new category of biodegradable polycation has been synthesized, delivering plasmids for a high biological effect.161–162 The polycation is based on grafted oligoamine residues of natural polysaccharides. The grafting means the side chain oligomers are attached to either a linear or branched hydrophilic polysaccharide backbone, allowing two- or three-dimensional interaction with an anionic surface area typical to the double- or single strand DNA chain. Low molecular weight cations and their lipid derivatives such as LipofectionTM and Lipofectamine® have a localized effect on the DNA, and the degree of complexation is dependent on arrangement of small molecules around the anionic DNA.163 Each molecule has to be synchronized with the other molecules during the transfection process, whereas the oligoamines grafted onto a polymer are already synchronized and each side chain helps the other side chain arrange to fit the anionic surface of a given DNA.164 The use of biodegradable polysaccharide carriers is suitable for transfection and biological applications because of their water soluble nature, readily transporting to cells in vivo by known biological processes, and acting as effective vehicles for transporting agents complexed with them.165

1.16

Chitosan

Chitosan is a biodegradable polysaccharide composed of two subunits, Dglucosamine and N-acetyl-D-glucosamine, linked together by b-(1, 4) © 2008, Woodhead Publishing Limited

28

Natural-based polymers for biomedical applications

glycosidic bonds. The unique physicochemical and biological properties of chitosan present it as a promising candidate for macromolecular delivery such as DNA and proteins. Positively charged amines of chitosan interact with negatively charged plasmids and condense them into compact structures. The low toxicity, biodegradability, and biocompatibility of chitosan make it a suitable candidate for gene delivery purposes.166–168 Electrostatic interaction between chitosan and DNA is strong so that the polyplex does not dissociate until it is delivered to the cell.169 A treatment of DNA/chitosan complex with phosphate buffered saline resulted in the release of only 0.05% of DNA, indicating a tight complex.170 Transfection efficacy of low molecular weight chitosan has also been examined.171 Chitosan was condensed with DNA above a 1:2 weight ratio (plasmid/ chitosan) (its formulation efficiently protects DNA from DNase degradation172) and the distribution of encapsulated DNA in chitosan nanoparticles was examined by ethidium bromide (EtBr) quenching assay in combination with confocal laser scanning microscopy. Another method for determining DNA loading is by PicoGreen® assay after digestion with chitosan and lysozome. Simply by the mixing of chitosan and DNA, 95% loading of DNA is achieved. Encapsulation yield is attributed to the shift of DNA conformation from a supercoiled state to a relaxed state, as judged by gel electrophoresis. Stability studies showed that cross-linked chitosan/DNA nanoparticles in water are stable for more than three months, whereas the uncross-linked formulation is stable only for several hours. 173 The characterization of chitosan-DNA nanoparticles, the evaluation of their transfection efficacy and their effect on cell viability on human osteosarcoma cells (MG63), MSCs, and HEK293 has been determined.174 High gene expression was achieved with HEK293 cells in vitro using chitosan as the gene carrier compared to MG63 and MCS cell types. Cell viability studies following incubation with nanoparticles confirmed the lack of toxicity of chitosan. Low cytotoxicity and the ability to transport and release genes intracellularly makes chitosan/DNA nanoparticles a potential candidate for nonviral gene delivery.

1.16.1 Chitosan derivatives Several chitosan derivatives have been synthesized in the last decade to obtain a modified carrier with altered physicochemical characteristics.175 The modification includes quaternarization of amino groups to increase the net positive charge of the complex, ligand attachment for targeting purposes, conjugation with hydrophilic polymers to increase the stability of the chitosan–plasmid complex against degrading enzymes, conjugation with endosomolytic peptide to increase the efficiency of transfection, etc.

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Deoxycholic acid modified chitosan Deoxycholic acid modified chitosan was prepared using ethylene dichloride (EDC) as the coupling agent in a methanol/water medium,147,175 and the degree of substitution was defined to be 5.1 (5.1 deoxycholic acid groups substituted per 100 anhydroglucose units). Hydrophobically modified chitosan provides colloidal stable self-aggregates in aqueous media, forming particles of approximately 160 nm diameter. Self-aggregate DNA complexes were prepared in aqueous media, used in transfecting mammalian cells in vitro. The transfection efficiency of this system was relatively high in comparison with naked DNA but significantly lower than the Lipofectamine®/DNA formulation. Quaternarized chitosan Another approach to increase the transfection rate using chitosan is the preparation of trimethylated chitosan oligomers (TMO) through quaternarization of oligomeric chitosan. This process is based on a reductive methylation procedure using methyl iodide in an alkaline environment. Trimethyl chitosan derivatives of 40% (TMO-40) and 50% (TMO-50) degrees of quaternarization were synthesized and examined for their transfection efficiencies in two cell lines: COS-1 and Caco-2.176–179 TMO-50 markedly increases the transfection efficiencies from 5-fold to 52-fold. TMO-40 displays even higher transfection efficiencies ranging from 26-fold to 131-fold. Chitosan and TMO oligomers were found to exhibit significantly lower cytotoxicity than DOTAP, a well known cationic lipid commonly used as a transfecting reagent. PEGylated chitosan Several methods have been developed for the grafting of hydrophilic polymers such as PEG onto chitosan to improve affinity to water or organic solvents.180–183 PEG–chitosan derivatives with various molecular weights (Mn = 550, 2000, 5000) of PEG and degrees of substitution were synthesized, and water solubility of these derivatives was evaluated at pH values of 4, 7.2, and 10.184 PEG modification was found to minimize aggregation and prolong the transfection potency at least for one month during storage. Intravenous injection of chitosan–DNA nanoparticles and PEGylated chitosan– DNA nanoparticles resulted in a majority of nanoparticles localizing in the kidney and liver within the first 15 minutes. The clearance of the PEGylated nanoparticles was slightly slower in comparison to non-PEGylated nanoparticles.

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Galactosylated chitosan Galactosylated chitosan-graft-dextran-DNA complexes were prepared and examined.185–186 Galactose groups were chemically bound to chitosan for liver targeted delivery, and dextran was grafted to enhance the complex stability in aqueous media. The system was found efficient to transfect liver cells expressing asialoglycoprotein receptor (ASGR), which specifically recognizes the galactose ligands on modified chitosan. Similarly, galactosylated chitosan-graft-PEG (GCP) was developed for the same purpose. GCP–DNA complexes were found to be stable due to hydrophilic PEG shielding and increased the protection against DNase. GCP–DNA complexes were found to enhance transfection in HepG2 cells having ASGR, indicating that galactosylated chitosan will be an effective hepatocyte-targeted gene carrier. The synthesis of lactosylated-modified chitosan derivatives has been done and their transfection efficiencies tested in several cell lines.187 However, in vitro the transfection was found to be cell-type dependent. HeLa cells were efficiently transfected by this modified carrier, even in the presence of 10% serum, but neither chitosan nor lactosylated chitosans have been able to transfect HepG2 and BNL CL2 cells. In vivo mediated transfection applying Dextran-spermine vector has been used as a gene carrier although cationic complexes have proved to be very efficient in transfecting cells in vitro. It has been well recognized that in vitro their effectiveness does not correlate with relatively poor activity in vivo. This low efficacy in vivo was attributed to the differences in the biology, functionality, and complexity between cell cultures and animal models as well as to the changes in the complexes’ structures upon their interaction with cells and biological fluids. Ideally, the complex should be delivered exclusively to target tissue, where it is subsequently taken up and further processed on the cellular level. However, in in vivo administration (i.v.), the complexes must go first through the biological milieu (a process that may include several obstacles). One of the limitations for potential transfection efficiency is rapid clearance of the polyplex from the blood circulation. For example, particles with strong positive charge associate with negatively charged biological membranes, blood proteins and lipoproteins to work as opponents. In vivo the study was done using dextranspermine (Fig. 1.11) as a nonviral self-assembled nucleic acid delivery system.188 The influence of the interaction with serum proteins on distance, surface potential, surface pH, complex biodistribution, and transgene expression using several routes of administration were characterized. It was demonstrated that local administration of polyplexes resulted in systemic distribution accompanied by transgene expression in the liver and lungs. In addition,

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high concentrations of polyplexes in the gastrointestinal tract indicated that polyplexes injected locally are transferred to distant organs through the blood circulation. This effect can be attributed to the much higher solubility of the polymer in aqueous media, combined with its lower positive charge, which makes its association with the cells at the site of administration weaker. This enables the polyplexes to flow with the blood to distant organs. This showed that the transgene expression in lungs is attributed to the high positive charge density and high NH3/DNA ratio, while involvement of the galactose receptor of the liver parenchymal cells is probably responsible for polyplex uptake in the liver. Systemic and local toxicity of dextran-spermine and its polyplexes was also studied. Polyplexes based on dextran-spermine were administered using the combined intranasal (i.n.) and intramuscular (i.m.) routes to increase levels of transgene expression and determine for expressed gene in the lung, liver, and muscle tissues. The transgene expression was dose- and charge ratio-dependent. Using the i.m. and i.n. routes of administration, the transfection takes place primarily in the bronchial epithelial cells, pneumocytes, and bronchial alveoli of the lungs, in the fibrocytes, and in the hepatocytes. Tissues which expressed the gene were further stained using hematoxylin and eosin and studied for local toxic effects. Systemic toxicity of dextranspermine and its polyplexes was also evaluated after i.m. administration. The following relevant parameters were examined in this study: the weights of animals, major organs (spleen, liver, lungs, heart) and blood analysis. Mild toxicity was revealed by histopathological assessment in the muscle and there were no abnormal findings in the liver or lungs. In addition, no systemic toxicity, no decrease in WBC counts, no thrombocytopenia, and no detectable increase in levels of serum transaminases were found.189

1.17

Applications of polysaccharides as drug carriers

As discussed previously, macromolecular polysaccharides and other natural as well as synthetic polymers offer potential applicabilities as high molecular weight carriers for various therapeutically active compounds.190,191 This section expands the discussion on the use of polysaccharides for improved drug delivery and targeting.

1.17.1 Dextrans Dextran is attached to the drug molecules for the formation of a prodrug using various techniques like direct linkage, attachment through intercalated spacer arm, use of modulator ligand and tissue specific receptor ligand.192 In the direct linkage model of dextran the drug is directly linked to dextran, which will release the active agent in a predictable manner. The regeneration

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O

O H O OH H H OH OH H H H OH Dextran

KIO4

OH

H

O H O OH

OH

H

OH

O OH

H

(m
m

O

Oxidized dextran

O

O

OH H Spermine

H

O

n

H

O

H

O H O OH

H OH

H

NaBH4

O

H H

N

m

O NH

O

OH

H N

Dextran-spermine (Imine conjugate)

NH2

H

H

OH

O

OH

H

CH2 NH

O

H

NH

H2C OH H N

m NH2

Dextran-spermine (Amine conjugate)

1.11 Grafting of spermine moieties on dextran. Dextran-spermine conjugates were synthesized by conjugation of spermine to oxidized dextran by reductive amination [Ref. 189b].

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of the parent drug would be exclusively governed by the pH dependent hydrolysis, as the bulky dextran matrix would be inaccessible to enzymatic attack. Intercalation of a spacer arm between the drug and the carrier dextran may serve three purposes. (1) The terminal functional group of the spacer arm was different which allows the covalent drug fixation to be established through a variety of chemical bonds. (2) Steric hindrance of enzyme activation of the liganded drug is circumvented by augmenting the distance between the drug and the dextran backbone. (3) Sequentially labile dextran prodrugs are constructed specifically, in which the pH dependent hydrolysis only liberates the spacer-drug derivatives, which after extravasation or diffusion from the site of injection is activated at the diseased tissue. The active drug is released from the dextran prodrug by the cleavage of the covalent bond existing between drug and the carrier moiety dextran, enzyme or pH. If the drug contains a hydrolyzable chemical bond, it might deteriorate even attached to the polymer backbone. The ability of the bulky dextran molecule to suppress catabolism of the attached drug/enzyme has been exploited extensively.193 Partial enzymatic disrupture of the main chains of dextran prodrugs changes the molecular weight distribution of the originally administered prodrug, making it susceptible to the attack of various hydrolases.194

1.18

Applications of dextran conjugates

1.18.1 Targeting tumor cells Methotrexate (MTX)–dextran conjugates have been synthesized by covalently linking MTX to dextran through a short-lived ester bond (MTX-ester-dextran) and a longer-lived amide bond (MTX-amide-dextran). The ability of these agents to kill cells and to penetrate through tissue has been evaluated. The cytotoxicity of MTX-ester-dextran and MTX-amide-dextran was found equivalent to unmodified MTX. Intracranial polymeric delivery of MTX or MTX-amide-dextran to rats with intracranial 9L gliosarcoma produced modest but significant increases in survival; conjugation of MTX to dextran appeared to shift the dose-response curve to a lower dosage.195 Synthesis of dextran conjugate of sodium phenylacetate (NaPA) with substituted dextrans like dextran-methyl-carboxylate-benzylamide (LS17DMCB) has been reported, for better tumor inhibition effect against human tumor melanoma 1205LU cells than NaPA alone.196 A dicarboxymethyldextran conjugate of cisplatin is synthesized by immobilizing cisplatin to dextran through a six-membered chelate type coordination bond, which shows longer half life and better tumor growth inhibitory activity than plain cisplatin in colon 26 cancer cells.197 A macromolecular prodrug of cisplatin using dextran carrier with branched galactose units has been synthesized for targeting to hepatoma cells effectively.118

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Clinically available camptothecins (CPTs) like irinotecan (CPT-11) and topotecan represent one of the most prominent classes of antitumor agents. A new macromolecular prodrug denoted T-0128 was synthesized bearing a novel CPT analog T-2513 conjugated to carboxymethyl dextran via a triglycine spacer to improve their pharmacological profile.198 This conjugate shows better specificity and ten times better activity than its parent T-2513. The conjugate of paclitaxel with carboxymethyl dextran via an amino acid linker produces better antitumor activity than paclitaxel alone.199 The pharmacokinetic and therapeutic studies of mitomycin-dextran conjugate shows better activity and prolonged action than plain mitomycin against walker-256-carcinoma in rats.200 The effect was distinctly bitter in case of cationic conjugate which is attributed for the presence of a high load of negatively charged sialic acid residues on cancer cell surface accomplishing effective cationic conjugate absorption.201 Over-expression of the epidermal growth factor receptor (EGFR) has been caused in bladder cancer and is a potential target for therapy with radionuclides. The intravesically administered EGF-dextran conjugate EGFdextran-99mTc selectively accumulates in the tumor tissue and shows better activity than 99mTc alone.202

1.18.2 Targeted antimycobacterial therapy Tuberculosis treatment requires a long-term antibiotic therapy. Targeted antibiotic therapy improves the efficacy of treatment by concentrating the drugs close to mycobacterium. Antibiotic norfloxacin has been linked to mannosylated dextran using a peptide spacer arm for this purpose. This conjugate shows more efficiency against mycobacterium than plain norfloxacin.203

1.18.3 Improved in vitro physico-chemical properties The tendency of enzymes to undergo autolytic degradation and thermal lability might be partly circumvented by coupling with dextran. The thermal stability of various enzymes like adenosine deaminase,204 a-amylase,205 epoxide hydrolase206 and a-chymotrypsin207 has been increased by their conjugation with dextrans of molecular weight 1 000 000, 80 000, 70 000 and 10 000 respectively. Possibly other bioactive peptidic agents might be stabilized in the same way. Several carboxylic acid drugs like NSAIDs are insoluble in water in the free acid form. Dextran prodrugs of such drugs may constitute an alternative approach to provide reproducible bioavailability of water insoluble drugs by releasing the active agent in a soluble state in the gastrointestinal tract.208 The water solubility of naproxen is increased by a factor of 500 when it is conjugated to dextran.209 Various conjugates of metronidazole have been prepared with dextran, like dextran metronidazole

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monomaleinate ester conjugate, dextran metronidazole monoglutarate ester conjugate, and dextran metronidazole monosuccinate ester conjugate in order to improve physico-chemical properties of metronidazole.210

1.18.4 Antifungal activity Life-threatening fungal infections have become increasingly widespread, especially among immunocompromised patients, such as those undergoing cancer treatment or transplantation and those with AIDS. The current treatment of severe systemic fungal infections is inadequate, due to limited availability of effective parenteral drug formulations and the appearance of new opportunistic fungal infections resistant to the marketed drugs. The pattern of new or newly resistant species emerging in response to widespread and prolonged drug treatment renders the development of new effective parenteral antifungal drug delivery systems highly important.211 Synthesis of dextran conjugate with Nystatin (Nys) has been reported for producing a stable water-soluble polysaccharide conjugate of Nys as antifungal treatment.212 Nys, a tetraene-diene antifungal agent, possesses a broad spectrum of activity. Problems associated with its insolubility in injectable solvents and high toxicity in iv injections have precluded its systemic administration.211 Nys is unstable in aqueous medium and is sensitive to oxidation and heat.213 The drug was conjugated by a Schiff-base reaction with oxidized dextran. High conjugation yield of active Nys was obtained. The conjugates were highly water soluble and could be appropriately formulated for injection. The conjugates showed comparable MIC values against Candida albicans and Cryptococcus neoformans and were about 25 times less toxic than free Nys after a single injection in mice.

1.18.5 Antileishmanial targeted therapy Pyrimethamine is effective in antileishmanial therapy because of its ability to inhibit both the enzymes of leishmanial folate pathway. But in vivo this effect occurs only at high concentration, associated with toxicity. To overcome this problem, a prodrug of pyrimethamine, namely carboxymethyldextran thiomannopyrannoside pyrimethamine, has been synthesized and utilized with approximately 50% destruction of intracellular amastigotes with no detectable toxicity to macrophage cells.214

1.18.6 pH-controlled intracellular drug release Polymers like dextran, when entering the endosomal or lysosomal compartment, are exposed to an acidic medium (pH 4.5–5.5). Using acid cleavable spacers like hydrazon spacer and N-cis-aconityl spacer between the drug and the

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carrier, a pH controlled intracellular drug release can be obtained.215 Streptomycin has been linked to dextran via a carboxylic hydrazone linkage for its intracellular delivery.216

1.18.7 Pharmacodynamic applications: Modification of immunogenecity of peptides When proteins or peptides are conjugated to macromolecules, the resulting conjugate colloid in most cases and the protein core get protected from interaction with other macromolecular plasma components by a hydrophilic polymer shield. Therefore the conjugation decreases the immunogenicity of the protein component and permits repeated administration of foreign proteins. This approach is useful when the protein is an enzyme with a low molecular weight substrate found in plasma. The antigenicity of L-asparaginase217 and catalase218 was substantially suppressed when they were conjugated to oxidized dextran and dextran with molecular weight of 40 000, respectively. Spacer linked dextran derivatives of pancreatic RNase showed decreased antigenicity.219

1.18.8 Pharmacokinetic applications: Systemic sustained drug action The profile of plasma concentration of drugs is an important determinant of their quantitative access to peripheral targets. The plasma concentration is usually measured as the area under the curve (AUC). In general, slow renal elimination and metabolic inactivation promote better access of drugs to remote targets. The efficacy of various drugs is limited due to their rapid renal excretion. Conjugation of drugs with hydrophilic macromolecular carriers like dextran prevents rapid renal excretion and restricts the drug entry into cells, thus prolonging their plasma circulation time.220 Insulin dextran complex when injected intraperitoneally in diabetized rabbits maintains blood glucose levels at normal value over several days.221 Design of dextran prodrugs prolonged systemic concentration of the liberated drug after oral administration. This has been reported for quinidine,222 ursodeoxycholic acid,223 ibuprofen224 and analogs of aspirin.225 Isoniazid has been conjugated to dextran to form dextran-isoniazid complex, which exhibits prolonged persistence in circulation and reduced acute toxicity in guinea pigs as compared to the parent drug.226 Nicotinic acid is an effective hypolipidemic agent, but suffers from the disadvantage of rapid elimination from the systemic circulation. When it is conjugated with dextran, the drug is slowly released from the polymeric support and prolonged the presence of the active substance in the body to lower the triglyceride levels.227

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1.18.9 Overcoming multidrug resistance A major problem of using antitumor agents in cancer chemotherapy is multidrug resistance. This occurs mainly due to over-expression of the P-glycoprotein (Pgp). This transmembrane glycoprotein, encoded by the mdrl gene, functions as an energy dependent efflux pump reducing the intracellular accumulation of anticancer drugs.228 When the free drugs are administered, they enter the cell by diffusion through the plasma membrane and are recognized by the Pgp pumps. On the other hand when these drugs are conjugated to macromolecular carriers, the drug in the form of polymeric prodrug is taken up by endocytosis and subsequently efflux pumps are circumvented in turn reducing multidrug resistance. The antitumor antibiotic doxorubicin was conjugated with polymeric dextrans of various molecular weights and the cytotoxicity of the conjugates against human carcinoma KB-3-1 cells and its multidrug-resistant subclone KB-V-1 cells was measured. The conjugates were found less toxic to the KB-3-1 cells than the free doxorubicin. Furthermore these conjugates could act synergistically with other cancer drugs.229

1.18.10 Reduction of toxicity The effects exerted on normal cells by antineoplastic agents needs the administration of relatively low doses of drugs.230 Improved chemotherapy might arise from drug delivery systems using polymers, providing the maintenance of prolonged moderate body levels of the drugs. In contrast to plain daunomycin, the dextran daunomycin conjugate does not show side effects like atrophy of spleen and bonemarrow, damage to the heart and liver.231 In comparison to parent daunomycin, the LD50 of the conjugate was enhanced by a factor of approximately 3.

1.18.11 Gastrointestinal relief Gastrointestinal side effects constitute the most adverse reactions of nonsteroidal antiinflammatory drugs. Flurbiprofen causes gastrointestinal disturbances, peptic ulceration and GIT bleeding. Conjugation of flurbiprofen with dextrans of different molecular weights (40 000, 60 000, 1 000 000, and 2 000 000) shows a remarkable improvement in physico-chemical properties, colon specificity and gastrointestinal side effects.232 Conjugates of suprofen have been synthesized by preparing their acylimidazol derivatives, which are condensed in situ with dextrans of different molecular weights (40 000, 60 000, 1 100 000 and 2 000 000). The ulcerogenic index of the conjugates shows a remarkable reduction as compared to the parent suprofen.218 The nephrotoxicity of gentamycin is substantially reduced when it is conjugated with dextran sulphate.

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1.19

Natural-based polymers for biomedical applications

Site-specific drug delivery

1.19.1 Colon-directed drug delivery Natural polysaccharides have been extensively used for the development of solid dosage forms for delivery of drugs to the colon, as it is inhabited by a large number and variety of bacteria and secretes many enzymes like β-Dglucosidase, β-D-galactosidase, amylase, pectinase, xylanase, β-D-xylosidase and dextranase.234 A formulation of drug–saccharide conjugate (prodrug) is an important approach for targeting drugs to the colon. Among the various polysaccharides, dextran is an important carrier for delivery of drugs to the colon.235 5-Aminosalicylic acid (5-ASA) is an effective drug for inflammatory bowel disease, but it is absorbed rapidly in the stomach and small intestine hence a negligible amount of drug only reaches the colon. This problem is solved by synthesizing a number of azo-coupled dextran-5-ASA prodrugs,236 which effectively deliver 5-ASA to the colon. Dextran-nalidixic acid ester (dextran-NA) was developed as a colon specific prodrug.237 When release studies were performed in HCl buffer (pH 1.2) and phosphate buffer (pH 6.8) at 37°C NA was not detected within 6 h of the incubation period, indicating that dextran-NA might be chemically stable during the transit through the gastrointestinal tract. When dextran-NA was incubated with cecal contents of rats at 37°C, the extent of NA released in 24 h was 41% of the dose. Dexamethasone-succinate-dextran (DSD) is synthesized by linking dexamethasone to dextran (Molecular mass 70 400 Dalton) using succinic anhydride as a spacer to deliver dexamethasone specifically to the colon.238– 239 Methylprednisolone has been covalently attached to dextran using succinic acid and glutaric acid as the linkers.240–241 The hydrolysis kinetic studies show that conjugation with dextran helps to deliver the drug to the large bowel.

1.20

Pectin drug site-specific delivery

Polysaccharide pectin is used for specific drug delivery. Several pectin formulations have been developed and tested in vitro, ex vivo, and in vivo for their ability to deliver bioactive substances for therapeutic purposes to the living tissues. Pectin derivatives possessing primary amino groups were found to be more mucoadhesive and are used in nasal and other mucosal drug delivery. Pectin derivatives with highly esterified galacturonic acid residues are found to be more hydrophobically capable of sustaining the release of incorporated fragrances for a prolonged time. Less esterified pectin derivatives penetrate deeper into the skin and are used in aromatherapy formulations. Pectin in combination with a corn protein zein forms hydrogel beads. This bounded zein restricts the bead swelling and retains the porosity of the beads. However, the pectin networks protect the zein from protease attack. These complex beads are used as ideal vehicles for colon-specific drug delivery.77

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1.20.1 Pectin nasal drug delivery Nasal route of drug administration is used as an alternative to injection. Nasal drug delivery was traditionally used for the treatment of local nasal diseases, such as congestion, allergy, infection and inflammation. Now, this route of drug administration is considered for systemic drug delivery. The entire nasal cavity, except for the vestibule of the nose, is lined with mucosal tissues, which functions in filtering out small airborne particles and warming, and humidifying inhaled air before the air reaches the lung. For these purposes, the nasal cavity is rich in blood vessels. The veins of the nasal mucosa form a venous plexus in the connective tissue of the nose. Due to the high degree of vascularization, the major advantage offered by nasal drug delivery is the high nasal permeability. It only takes about 15 min for a vaccine to transport through the nasal mucosa into the circulatory system.242 In addition to rapid absorption, advantages of nasal drug administration include high patient compliance, self-medication, mild environmental conditions, no exposure to acid/basic pH solutions and no protease attack, and the avoidance of liver metabolism.243–245 The limitations of nasal drug delivery are mainly due to the small absorption area in the nasal cavity and short residence time. The maximum dose to be given is normally less than 150 µl and the dosage formulations are quickly cleared. Thus, low water soluble drugs need adequate formulations for sufficient absorption by nasal mucosa. In therapy that demands constant drug blood level, prolonged drug release formulations are required; otherwise, frequent nasal administration is required. Pectin-based formulations are expected to be able to meet some of these challenges by facilitating drug adsorption and bioavailability in nasal cavities, and serve as an effective and ideal nasal drug delivery system. Pectin macromolecules diffuse into nasal mucosal tissues and pectin gel formulation regulates the adsorption of incorporated drugs.246 Pectin adsorption in nasal tissues was dependent on gel concentration and the side chain functional groups of the pectin. Another important advantage of pectin formulations is the exclusion of the use of organic solvents in dosage preparation. This biopolymer is a wellknown gelling material and gels in aqueous solutions under mild conditions. Pectin derivatives also coacervate easily with calcium, proteins, polypeptides and polyphosphates. A large number of drugs incorporate into pectin formulations by various physical methods, such as diffusion, mixing, encapsulation or co-precipitation. These lead to the avoidance of drug denaturation and possible irritation of the dosage forms to mucosal tissues. Biocompatibility has been examined in a rat model using pectin formulations without drug. No adverse effects of intranasal administration in the form of gels or microspheres were identifed by visual observation or histological examinations.

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1.20.2 Pectin oral drug delivery Pectin derivatives have been used as raw materials to construct drug carriers for oral drug delivery. Several groups have worked on this possibility.234,247–248 Polysaccharide pectin has a long and safe use history in the food industry. Furthermore, pectin passes intact through the gastrointestinal tract and is degraded by colonic microflora, which remains relatively consistent across diverse human populations. It focuses on the use of pectin as a drug carrier for colon specific drug delivery. One disadvantage of pectin for colonspecific drug delivery is that pectin formulations swell considerably in physiological conditions. Drugs with high solubility in physiological conditions may display a premature release due to the expanded pore size of pectin formulations. This can be avoided by using pectin in combination with other polymers to form more stable structures.249–251 Several pectin composites have already been developed for this purpose. A pectin/zein complex hydrogel bead is one of these composites. The pectin/zein complex hydrogel beads were prepared by pumping pectin gel into an ethanol solution containing zein and calcium chloride at designed ratios. The zein was mainly located around the periphery of the beads to form a shell-like structure, but also migrated into the beads, where the protein macromolecules were bound to the pectin networks or aligned to densely packed fibers. Dissolution experiments showed an improved stability of the complex beads in environments that mimic the upper GI tract. The inclusion of a small portion of zein into the pectin networks not only suppressed the swelling behavior of pectin, but also offered the beads great protease resistance. The ‘shell’ around the beads restricted the swelling of pectin networks; in turn, the steric hindrance of the pectin networks shielded the bound zein from protease digestion. Meanwhile, the in vitro study showed no effect of zein coating on pectinase-induced degradation of beads in a medium mimicking the lower GI tract. In practice, a single dose of complex mixture beads with different zein contents was used to deliver drugs at different time intervals. These systems may be useful for oral vaccine or antibiotic delivery. Oral vaccine delivery requires stimulation of antibody producing cells at two or three successive time points within a day or two. A single administration of a cocktail of different pectin/zein beads is expected to provide such stimulation. An example is to deliver antibiotics such as cipro oxacin against Bacillus anthraces or anthracis. For this case, a single administration of a cocktail of selective pectin and pectin/zein beads provide a sufficient high serum level of the drug for a prolonged time.252

1.21

Liposomal drug delivery

Drug delivery with liposomes as carrier systems provide opportunities for designing site-specific drug therapy. The engineered or tailored versions of © 2008, Woodhead Publishing Limited

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liposomes offer potentials of exquisite levels of specificity and drug targetability. The macromolecular polysaccharides are used as molecular carriers such as dextran, mannan, amylopectin and pullulan, either in their native form or as carrier-conjugates. Naturally occurring polysaccharides protected the cell plasma membranes against physico-chemical stimuli, such as osmotic pressure and ionic stress. However, on adsorption to lipid carriers, peptization or coagulation of the system may occur; thus, partially hydrophobized polysaccharides are used in drug delivery.253 Polysaccharides are used in liposomal coating because of their biodegradability, protein rejection ability, low toxicity and interactions in biological process through sugar moieties.254– 255 Naturally occurring polysaccharides are used as vesicle coatings; polysaccharides are targeted to the surface of macrophages and the brain tumor cells.256 However, in particular modifications, dextrans exhibit heparin like properties such as anticoagulation capacity, inhibition of smooth muscle cell growth, antiviral activity, etc. Dextran was used in this study because dextranpolysaccharide stabilization of the liposomes increases circulatory half-life and reduces protein absorption. The uptake of such liposomes suggests that the functionalized dextran coated liposomes are used as biomaterials for the stable carriers of drugs in a systemic administration.257 Hirudin is a potent and specific inhibitor of thrombin entrapped liposomes with a palmitoyl dextran coating.258 Palmitoyl dextran coated liposomes were found to stabilize hirudin and result in greater retention of hirudin’s ability to inhibit thrombin’s enzymatic activity. Oral delivery of potential palmitoyl pullulan (OPPu) and cholesteroyl pullulan (CHPu) coated liposomes against the challenges of detergent and bile freeze-thaw cycling and long-term storage has been also investigated.259 No significant changes in the vesicle size, integrity and drug content were observed and both OPPu and CHPu coated liposomes exhibited exceptional stability as compared with plain formulations.

1.21.1 Therapeutic and clinical applications Recently, various polysaccharides and their derivatives have been used for drug delivery and targeting strategies. Some of the therapeutic applications of polysaccharide vesicles are discussed. Lung therapeutics Polysaccharide liposomes are known for their selective and potential drug delivery systems for the therapy of lung diseases.260 O-palmitoyl amylopectin (OPA) liposomes are reported to be sequestered and retained selectively in lungs by anionic-scavenging receptors. Liposomes appended with O-palmitoyl pullulan (OPP) and O-palmitoyl amylopectin (OPA) are rapidly cleared from

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the blood as compared to ‘naked’ liposomes. Though they have a relatively wide tissue distribution including liver and spleen, it is found that OPA liposomes are selectively intercepted, sequestered and internalized by the lung macrophages and monocytes. Subsequent to this observation, investigations were made on OPA anchored liposomes to explore their potential as a delivery system for sisomycin treatment of lung diseases in guinea pigs infected with Legionella pneumophila.261 The therapeutically beneficial results of these studies were subsequently promoted. Specifically, amylopectin liposomes were found to be effective for the delivery of ceftazidine to L pneumophila infected guinea pigs to treat with the free drug; the survival rate achieved following the liposome treatment was 30%. Liposomes with amylopectin show good efficacy as compared to liposomes conjugated with PEG in terms of improved lung uptake after intravenous infusion. Amylopectin encapsulation anchored phosphatidylethanolamine (PE) and cholesterol (70:10:20 mol %) liposomes with dextran (Dx) or functionalized dextran (FDx) are hydrophobized to penetrate the lipid bilayer during vesicle formation. The study was performed using radiolabeled markers and fluorescent probes and revealed that coating of liposomes with FDx enables specific interactions with human endothelial cells in culture. Conclusively, bioactive polymer based liposomes hold promise as an attractive approach for vascular cell targeting. Targeted chemotherapy The polysaccharide-based liposomes have been used as stable and targetable drug carriers in effective chemotherapy, introducing chemo-therapeutics into target cells or tumor cell lines. These systems possess a unique targetability to specific tissues such as alveolar macrophages and other macrophages of RES. Polysaccharide-based liposomes could be employed as carriers to construct side-specific sensing molecule(s) such as MoAb to work against the tumor surface antigen, attached physically or chemically.265 The system being used as site-specific transport of sufficient quantity of an anti-tumor drug, release it within the vicinity of the target. Adriamycin bearing liposomes based on immuno-polysaccharide derivative demonstrated in-vivo targetability against human lung cancer (PC-9 grafted) in experimental athymic mice.262–263 Various polysaccharide based liposomals have been developed and engineered for brain targeting at human glioma.264– 265 They employed sulfatides and MoAb as site directing devices to endow targetability to the liposomes. Targeted chemotherapy of brain tumors using polysaccharide-based liposomes loaded with the antitumor drug cisplatin has been attempted.266 Recently, the polysaccharide-based liposomes have been exploited as carrier cargo equipped with a targeting ligand, where the polysaccharide coating

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stabilizes the system both in vivo and in vitro. Magnetoliposomes were used for hyperthermia based treatment of cancer.267 The liposomes were anchored with hydrazide pullulan to stabilize the phospholipid capsules and to provide an anchor for the immobilization of antibodies. When the antibody-conjugated and polysaccharide stabilized magnetoliposomes were incubated with cancer cells, they bond to the cell surface, being taken up by cells in about 12 times higher magnitude than the control formulations. Anti-tumor effects of polysaccharide-stabilized and ligand based liposomal Adriamycin on A66 hepatoma transplanted in nude mice was evaluated.268– 269 Tumor recognition ligand, 1-amino lactose (1-AL) was appended on the surface of cholesteroyl amylopectin (CHP)-based liposomes. The study was aimed at evaluating the role of polysaccharide coating on stability and tumor recognition ligand as a target site ligand. The uptake of these liposomes by AH66 rat hepatoma cells was estimated to be higher than liposomes without 1-aminolactose in vitro. Furthermore, 1-AL/CHP liposomal Adriamycin showed a stronger antitumor effect compared to other types of liposomal Adriamycin in vitro. These observations suggest that along with polysaccharide based liposomes, anticancer drug carriers are engineered for the active targeting to tumor cells. Systemic and mucosal vaccination Natural or bacterial origin polysaccharides show excellent immune response in association with protein carriers. Immunization potential of the bacterial polysaccharides encapsulated within liposomes is frequently suggested.270– 271 Recently, the potential of natural polysaccharide-based liposomes as an adjuvant for cell-mediated immunity has also been explored.272–273 The conjugate of cell wall protein and polysaccharide incorporated in a liposomal system may thus be presented as a potential adjuvant for oral vaccination against S. mutans vis à vis dental caries. A peptide-based vaccine, induced cell-mediated immunity. A 20-mer synthetic peptide, spanning the 98–117 amino acids of bovine leukaemia virus (BLV) envelope glycoprotein (Env) gp51 encapsulated in mannanbased liposomes has been used.274 Promising results were recorded when cDNA of HIV-1 was incorporated into mannan-based liposomes. The adjuvanticity of mannan-based liposomes for human immunodeficiency virus type-1 (HIV-1) DNA vaccine and the mechanism involved in the immunogenicity enhancement was studied. Coating of cationic liposomes with mannan significantly potentiated the vaccine and induced an HIV-specific delayed-type hypersensitivity (DTH) response. HIV-specific cytotoxic Tcell (CTL) activity elicited by DNA vaccination was also significantly potentiated on co-administration with mannan-liposome in the form of a therapeutic cocktail. This mannan-liposome-mediated activity was inhibited

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noticeably by pre-injection of anti-interferon (IFN) antibody, suggesting an important role of IFN in HIV-specific immune response. The results of both isotype-specific antibody and cytokine analysis revealed that mannan-based liposome DNA vaccination could prove to be a valuable tool for the enhancement of HIV-1 specific cell-mediated immune response.275–276 Macrophage activation The immuno-modulators themselves do not show any specificity or affinity to activate macrophages. Liposomes interact efficiently with macrophages; thus macrophages may serve as antigen presenting cells for liposomal antigens/ immuno-modulators. In order to modulate the in-vivo activity of several immunomodulators, they are generally administered encapsulated in polysaccharide-based liposomes. Polyanion polymers and synthetic polynucleotides were encapsulated into macrophage specific liposomes such as those anchored with mannan-Chol dervatives.277–279 Immuno-modulator activities of polysaccharide-based liposome possessing synthetic polynucleotide, polyvinyladenine and vinyladenine-alt-maleic acid (poly VAMA) have been evaluated.280 These studies shows macrophage activation and increased immunopotentiation of liposome encapsulated contents on surface of polysaccharides. Gastric mucoadhesion Among the polysaccharides used in drug delivery, chitosan has been widely used due to its bioadhesive properties. This is due in part to its characteristics of high molecular weight and degree of de-acetylation, its gel forming at low pH and also due to its poly-cationic character, which imparts its ability to bind strongly to mammalian cells. Chitosan based liposomes showed the highest percentage adhesion among the polymer-based liposomes. The mucoadhesiveness of chitosan-based liposomes was subsequently evaluated to develop a novel drug carrier system for oral administration of poorly absorbed drugs such as peptides. Muco-adhesive chitosan-based liposomes were used to improve oral absorption of insulin.281 After in vivo administration of the chitosan-based liposomes to male wister rats, the hypoglycemic response was prolonged over a period of up to 12 h. This sustained effect was attributed to the muco-adhesiveness of the system leading to an increased duration of contact with intestinal mucosa and hence an increased probability of insulin absorption. These studies explored the possibilities of polysaccharide based liposomes for the administration and muco-adhesion of poorly adsorbed drugs and macromolecules.

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195 Dang W, Colvin H B and Saltzman W M, Cancer Res, 1994, 54, 1729. 196 Gervelas C, Avramoglou T, Crepen M and Jozefonvicz J, Anti-Cancer Drug, 2002, 13, 37. 197 Ichinose K, Tomiyama N, Nakashima M, Ohya Y, Ichikawa M, Ouchi T and Kanematsu T, Anti-Cancer Drug, 2000, 11, 33. 198 Sezaki H and Hashida M, Crit Rev Ther Drug Carr Syst, 1984, 1, 1. 199 Sugahara S, Kajiki M, Kuriyama H and Kobayashi T R, Biol Pharm Bull, 2002, 25, 632. 200 Nomura T, Saikawa A, Morita S, Sakaeda K T, Yamashita F, Honda K, Takakura Y and Hashida M, J Control Release, 1998, 52, 239. 201 Marre A, Soyez H and Schacht E, J Control Release, 1994, 32, 129. 202 Bue P, Halmberg A R, Marquez M, Westlin J E, Nilsons S and Malmstrom P U, Eur Urol, 2000, 38, 584. 203 Roseeuw E, Coessens V, Balazuc A M, Micheline L, Chavarot P, Pessina A, Neri M G, Schacht E, Marchal G and Domurado D, Antimicrob Agents Chemother, 2003, 47, 3435. 204 Rosemeyer H, Kornig E and Seela F, Eur J Biochem, 1982, 122, 375. 205 Marshall J J, Carbohydr Res, 1976, 49, 389. 206 Ibrahim M, Hubert P, Dellacherie E, Magdalou J, Muller J and Siest G, Enz. Microb Technol, 1985, 7, 66. 207 Vegarud G and Christensen T B, Biotechnol Bioeng, 1975, 17, 391. 208 Fini A, Zecchi V and Tartarini A, Pharm Acta Helv, 1985, 60, 58. 209 Harboe E, Johansen M and Larsen C, Science, 1988, 16, 73. 210 Sternberg S, Science, 1994, 266, 1632–1634. 211 Domb A J, Linden G, Polacheck I and Benita S, Journal of Polymer Science: Part A: Polymer Chemistry, 1996, 34(7), 1229–1236. 212 Trachtenberg D M, Rodionovskaya E I, Gordina Z V, Rostovtseva L I, Kleiner G I and Nagle A M, Antibiotiki, 1966, 11, 9–13. 213 Larsen C and Johansen M, Int J Pharm, 1987, 35, 39. 214 Carvalho P B, Ramos D C, Cotrim P C and Ferreira E I, J Pharm Sci, 2003, 92, 2109. 215 Kratz F, Beye, U and Schutte M T, Crit Rev Ther Drug Carr Syst, 1999, 16, 245. 216 Coessens V, Schacht E and Domurado D, J Control Release, 1996, 38, 141. 217 Wileman T, Bennett M and Lilleymann J, J Pharm Pharmacol, 1983, 35, 762. 218 Marshall J J and Humphreys J D, J Appl Biochem, 1979, 1, 88. 219 Kurinenko B M, Kashkin A P, Kalacheva N V, Neringovo O F and Nekhoroshkova V, Biochem USSR, 1985, 50, 488. 220 Maeda H, Seymour L W and Miyamoto Y, BioconJ Chem, 1992, 3, 351. 221 Singh M, Vasudevan P, Ray A R and Guha S K, Macromol Chem, 1980, 181, 2433. 222 Astinotti D, Lapicque F and Dellacherie E, Macromol Chem, 1985, 186, 933. 223 Ghedini N, Ferruti P, Andrisano V and Scapini G, Synth Commun 1983, 13, 707. 224 Cecchi R, Rusconi L, Tanzi M C and Dannuso F, J Med Chem, 1981, 24, 622. 225 Grimova J and Hrabak F, J Biomed Mater Res, 1978, 12, 525. 226 Flip C, Ungureanu D, Gheorghita N, Mocanu G and Nechifor M, Rev Med Chir Soc Med Natl Iasi, 2003, 107, 179. 227 Moscow J A and Cowan K H, J Natl Cancer, 1988, 80, 14. 228 Lam W, Leung C H, Chan H L and Fong W F, Anti-Cancer Drug, 2000, 11, 377. 229 Arnon R and Sela M, Immunol Rev, 1982, 62, 5. 230 Hurwitz E, Wilchek M and Pitha J, J Appl Biochem, 1980, 2, 25.

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264 Yagi N, Naoi M, Sasaki H, Abe H, Konishi H and Arichi S, J Appl Biochem, 1982, 4, 121. 265 Yagi K, Naoi M, Glycolipid insertion into liposomes for their targeting to specific organs, in: K Yagi (ed.) Medical Applications of Liposomes, Tokyo, Japan Scientific Societies Press, 1986, 91. 266 Ochi A, Shibata S, Mori K, Sato T and Sunamoto J, Drug Delivery Syst, 1990, 5, 261. 267 Shinkai M, Suzuki M, Iijima S and Kobayashi T, Biotechnol Appl Biochem, 1995, 21, 125. 268 Ichinose K, Yamamoto M, Khoji T, Ishii N, Sunamoto J and Kanematsu T, Anticancer Res, 1998, 18 (1A), 401. 269 Yamamoto M, Ichinose K, Ishii N, Khoji T, Akiyoshi K, Moriguchi N, Sunamoto J and Kanematsu T, Oncol Rep, 2000, 7, 107. 270 Noguchi Y, Tateno M, Kondo N, Yoshiki T, Shida H, Nakayama E, and Shiku H, J Immunol, 1989, 143, 3737. 271 Noguchi Y, Noguchi T, Sato T, Yokoo Y, Itoh S, Yoshida M, Yoshiki T, Akiyoshi K, Sunamoto J and Nakayama E, J Immunol, 1991, 146, 3599. 272 Abraham E, Vaccine, 1992, 10, 461. 273 Sugimoto M, Ohishi K, Fukasawa M, Shikata K, Kawai H, Itakura H, Hatanaka M, Sakakibara R, Ishiguro M, Nakata M and Mizuochi T, FEBS Lett, 1995, 363, 53. 274 Ohishi K, Kabeya H, Amanuma H and Onuma M, Vaccine, 1996, 14, 1143. 275 Toda S, Ishii N, Okada E, Kusakabe K I, Arai H, Hamajima K, Gorai I, Nishioka K and Okuda K, Immunology, 1997, 92, 111. 276 Sasaki S, Fukushima J, Arai H, Kusakabe K I, Hamajima K, Ishii N, Hirahara F, Okuda K, Kawamoto S, Ruysschaert J M, Vandenbranden M, Wahren B and Okuda K, Eur J Immunol, 1997, 27, 3121. 277 Sato T, Kojima K, Iida T, Sunamoto J and Ottenbrite R M, J Bioactive Compatible Polym, 1986, 1, 448. 278 Oka M, Acta Med Nagasaki, 1989, 34, 88. 279 Sunamoto J and Sato T, J Chem Soc Jpn, 1989b, 161, 24. 280 Akashi M, Iwasaki H, Miyauchi N, Sato T, Sunamoto J and Takemoto K, J Bioactive Compatible Polym, 1989, 4, 124. 281 Takeuchi H, Yamamoto H, Niwa T, Hino T and Kawashima Y, Pharm Res, 1996, 13, 896.

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2 Purification of naturally occurring biomaterials M. N. G U P T A, Indian Institute of Technology Delhi, India

2.1

Introduction

The purification of naturally occurring biomaterials in early days was restricted to isolation and purification of so called natural products (Finar, 1964). The area was dominated by organic chemists who were using crystallization and distillation to purify plant extracts (Weissberger, 1965). The advent of chromatographic techniques using materials like silica and alumina created much excitement as the resolving power of these techniques was far superior (Heftmann, 1974). Tannins, alkaloids and terpenoids are some of the important classes whose purification yielded very useful information (and continue to do so). As biochemistry and microbiology developed, carbohydrates, proteins, lipids, nucleic acids became attractive target classes of molecules for separation. Microbial fermentation became a very important starting point for obtaining diverse classes of compounds which included vitamins and antibiotics (Verrall, 1996). Recombinant organisms, once the technology became available, became a favorite source of biomaterials especially proteins (Sambrook and Russel, 2001). Metabolic Engineering is also now emerging as a powerful a tool to produce secondary metabolites (Best, 1988; Glazer and Nikaido, 1995). Tissue culture techniques also made it possible to obtain a wide variety of products. The result was that with so many upstream strategies available, the classical purification approaches were perceived to be slow, inefficient and on the whole inadequate. Many new versions of chromatography were soon developed. Today, purification (of naturally occurring biomaterials) is one of the most complex areas in chemical and life sciences. This chapter gives an overview of what is available to people wishing to purify a particular class of compound.

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2.2

Classes of naturally occurring biomaterials

2.2.1

Cellular structures of living systems

55

All naturally occurring biomaterials, directly or indirectly, are obtained from cells of living organisms. Hence, a brief look at the cellular structure of living systems may be useful. Of the five kingdoms of living systems (animals, plants, fungi, protista and prokaryotes), the first four are collectively called eukaryotes (Heritage et al., 1996)

2.2.2

Constituents of cells

The cellular structures (Heritage et al., 1996) can be of two kinds. Prokaryotic cells do not have a membrane bound nucleus (the nuclear material is present as naked DNA), do not have membrane bound organelles in cytoplasm, have cell walls (mostly peptidoglycans but some other polymeric structures may also be there), have no chloroplasts or mitochondria and have small ribosomes. Eukaryotes share the property of having a membrane bound nucleus, DNA is complexed with proteins and present as chromosomes, their cytoplasmic organelles are membrane bound. The eukaryotic cells are 1000–10 000 times larger than prokaryotic cells.

2.2.3

Carbohydrates, proteins, lipids, nucleic acids as major biological molecules

In spite of considerable biodiversity of organisms, the key metabolic pathways of all cells surprisingly have similar designs (Alberts et al., 1994). Carbohydrates, proteins, lipids and nucleic acids play important roles in cell functioning. Also, cells do not function in isolation. Both inter- and intracellular communications play an important role in biology. Again, these classes of biological molecules play a key role in these processes and living organisms (Alberts et al., 1994). Carbohydrates Carbohydrates are classified as monosaccharides (e.g. glucose, fructose, mannose), disaccharides (e.g. lactose, sucrose) and polysaccharides (e.g. cellulose, starch, glycogen and chitin). The breakdown of glucose was one of the earliest metabolic pathways discovered. Simultaneously, its fermentation to produce ethanol turned out to follow similar molecular transformations. Its aerobic oxidation to CO2 and H2O constitutes an important energy yielding pathway in aerobic organisms. Starch (in plants) and glycogen (in animals) are storage molecules which breakdown to glucose when cellular dynamics require (Berg et al., 2002). Cellulose is the most abundant organic compound © 2008, Woodhead Publishing Limited

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on earth but is mostly present as lignocellulose material. It is estimated that about half of all the matter produced through photosynthesis is lignocellulose. As this material can be degraded in nature (and in the fermentors), lignocellulosic material constitutes the most important renewable resource available to the mankind (Glazer and Nikaido, 1995). In the quest for a sustainable society, it is believed that all chemical intermediates (which at present are mostly obtained from petroleum based processes) can, in principle, be obtained from lignocellulosic material. This concept of biorefinery embodies a major direction for sustainable industrial practices (Gupta and Raghava, 2007). Yet another important function of carbohydrates originates from more complex structures when carbohydrates conjugate with proteins (glycoproteins, proteoglycans) and lipids (glycolipids). The oligosaccharide chains decorating a protein molecule serve many important biological functions (Berg et al., 2002). Polysaccharides are industrially important molecules. Some industrially important bacterial polysaccharides are xanthan, dextran, alginate, curdlan and gellan. Scleroglycan and pullulan are fungal polysaccharides which are used in the industry. Major polysaccharides in order of decreasing consumption are cornstarch (and their derivatives), cellulose (and their derivatives), guar gum, gum arabic, xanthan, alginate, pectin, carrageenans, locust bean gum and gum ghatti (Glazer and Nekaido, 1995) Proteins Proteins (Berg et al., 2002) are made up of amino acids joined together by an amide bond (or peptide bond). All enzymes are proteins. Many important hormones are polypeptides. Lectins which have the property of specially recognizing carbohydrate residues are proteins. Antibodies, the key molecules of the immune system are also proteins. Based upon their biological functions, proteins can be classified as enzymes, regulatory proteins (e.g. insulin, interferons) transport proteins (e.g. hemoglobin, maltose binding protein), protective proteins (antibodies, thrombin), toxins, storage proteins, contractile proteins, structural proteins (e.g. collagen and elastin). Thus it is not surprising that purification of proteins from diverse sources is one of the most important areas in biotechnology. Proteins also occur in combination with other biological molecules (glycoproteins, lipoproteins, nucleoproteins). Conjugated proteins also include proteins containing heme or prosthetic group FAD. Lipids Structurally, lipids (Berg et al., 2002) constitute the most heterogeneous group of compounds which are put together on the basis of their property of

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dissolving in organic solvents rather than water. Fatty acids and fats (or oils), phospholipids, vitamins (fat soluble ones), steroids, terpenoids constitute this diverse class. Plasma membranes present in eukaryotic cells are bilayers of phospholipids associated with membrane proteins. On a weight basis, fat (or oil) gives twice as much energy as carbohydrate or protein during metabolic breakdown. While fats and oils (the triglycerides) are important as cooking and baking media and as salad dressings, monoglycerides (and diglycerides) find diverse application as biosurfactants (Gupta, 1996). Biodiesel, an important biofuel is an alkyl ester of long chain fatty acids and is obtained by transesterification of fatty acids (Shah and Gupta, 2007). Nucleic acids Either in the naked form or in association with proteins, nucleic acids (Davidson, 1965; Berg et al., 2002) constitute genetic material. DNA isolation and manipulation is at the heart of molecular biology. Plasmids are small DNAs which are present in many bacteria and eukaryotic organisms and are important in the context of recombinant DNA technology. The length of a nucleic acid is generally expressed in terms of kilobases; 1 kb is roughly equal to 660 kDa. In rapidly growing E coli cells, merely 0.2% of the total cell weight is accounted for by ‘other small molecules’ which may include heme, quinones, and other metabolites. Water constitutes 70% of weight and 1% consists of inorganic ions. The rest consist of carbohydrates, proteins, lipids and nucleic acids and their precursors (Bailey and Ollis, 1986).

2.3

Downstream processing of small molecular weight natural products

Globally, the chemical industry (excluding the pharmaceutical industry) is reported to have a turnover of about 71300 billion. The pharmaceutical industry would account for another 7541 billion (Jenck et al., 2004). While a good percentage of chemical products are synthetic, many are isolated from plants or fermentation. Antibiotics, vitamins, amino acids, drugs, fragrances and flavoring compounds are among the major classes. Any product recovery operation essentially consists of: (a) removal of particulates by filtration, centrifugation, sedimentation/decantation; (b) primary isolation by solvent extraction, adsorption, precipitation and ultrafiltration; (c) Purification by fractional precipitation, chromatography and crystallization (Verrall, 1996; Bailey and Ollis, 1986). Often, multiple unit operations utilizing a variety of techniques are required. Scale up considerations and economics play an important role in deciding these downstream operations. Lately, the ‘greenness’ of the process has also become a major consideration.

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2.3.1

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Solvent extraction

Liquid-liquid extraction is an important technique for obtaining a large number of natural products from fermentation broths (Verrall, 1996). Some of the variations in terms of mechanisms are physical extractions, ion-pair extraction and liquid ion exchange (Thornton, 1992). Examples of products which are obtained by solvent extraction are ethanol, lactic acid, citric acid, penicillin G, cephalosporin, streptomycin, vitamin B and riboflavin (Verrall, 1996). Essentially, solvent is so chosen so as to facilitate selective and high partitioning of the desired product in its phase. Generally, a pretreatment step, (generally filtration or centrifugation) is required for removing solid biomass. This generally leads to some loss and whenever possible, direct extraction of this product from the broth (‘whole broth’ extraction) should be attempted. Unfortunately many surface active molecules in the solid biomass limit interfacial mass transfer kinetics and also sometimes prevent separation of phases. Efficient contact between the product in the broth and the extracting solvent is the key parameter. Both column contactors (e.g. Karr column or reciprocating Karr column) and centrifugal contactors (e.g. Podbielnak contactor, extraction decanter) have been described (Verrall, 1996).

2.3.2

Supercritical fluid extraction

Fluids above their critical temperature and pressure are called supercritical fluids. CO2 has been the most widely used supercritical fluid for extraction in the food industry and its use in other sectors such as pharmaceuticals is being investigated (Verrall, 1996). The cost of the technique has prevented its wider applications. Its environmentally friendly nature has been its great attraction.

2.3.3

Membrane based methods

Size difference is the main criteria by which membranes separate molecules. For molecules with diameters in the range of 10 Å–1000 Å, ultrafiltration can be used for both concentration and separation (Bailey and Ollis, 1986). Composite membranes are used for better retention and consistent performance. The membrane selection is based upon molecular weight cut-off, robustness, low fouling property and cleaning – in-place possibility (Lutz and Raghunath, 2007). With the availability of commercial membranes with good performance characteristics, membrane based methods are increasingly becoming popular especially in the pharmaceutical industry. The various membrane based reactor designs are available (Lutz and Raghunath, 2007). Some special approaches using membranes are described below.

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Smart membranes Smart hydrogels take up solvent as a result of specific small changes in the environment (called stimulus). The swelling is reversible and can be a few hundred times the increase in volume (Roy and Gupta, 2003; see also Chapter 5 of this volume). Temperature sensitive hydrogel membranes have been described for size based separation of molecules (Feil et al., 1991; Nonaka et al., 1994). With time, this is another approach which should see some interesting applications in the purification of natural products. Pervaporation Pervaporation consists of a selective sorption of a liquid on a membrane, its diffusion through a membrane and its conversion to the vapor phase on the other side of the membrane due to application of a vacuum. It is an energy intensive process which can separate liquids (Fleming, 1992). Selectivity, permitting high diffusivity and mechanical strength are the three features which are important in designing the membrane. Many composite membranes have been described (Ruckenstein and Sun, 1995; Hirotsu, 1987). Pervaporation may in suitable cases (where liquid natural products are involved) prove valuable provided economics permit and a suitable membrane can be fabricated.

2.3.4

Chromatographic methods

Chromatographic methods are indispensable where high purity of the isolated product is desired. Adsorption chromatography on silica and alumina has been practiced by organic chemists for many decades. Non-ionic synthetic adsorbents like polystyrene resins and polymethacrylate resins can also be used for a wide variety of applications (Varrell, 1996). Ion exchange chromatography, over the years, has evolved into a very versatile technique. Not only a very large variety of matrices (with distinct capacity, pore size, etc.) have become available, different elution procedures such as displacement have been developed (Varrell, 1996). Hostettmann and Marston (1997) have described the applications of these preparative chromatography techniques to the purification of natural products. The book also discusses the separation of chiral molecules.

2.3.5

Imprinted polymers

Molecular imprinting consists of creating tailor made recognition sites in a polymer (Anderson et al., 1994; Alvarez-Lorenzo and Concheiro, 2006). A wide range of imprinted polymers including imprinted hydrogel have been

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described for a variety of small molecular weight substances (Chen et al., 2002).

2.3.6

Superamolecular assemblies

The discovery of host–guest complexation chemistry with crown ethers has led to a vast area which is often called molecular recognition (Frost, 1999). Cryptates, Lariat crown ethers, bis-crown ethers, paracyclophanes, fullerenes, cyclodextrins, cavitands, spherands, callixarenes and micelles is an illustrative list of such materials. In many cases, practical applications in separations are very few. The focus so far seems to be on looking at these designs as biomemetics.

2.4

Purification strategies for proteins

2.4.1

Post-genomic era

Earlier, proteins were obtained from naturally occurring microbial plant and animal sources. If one wanted a protein with a particular biological activity (e.g. enzyme, hormone or lectin), one had to look for an appropriate source which had an adequate amount of that biologically active protein (Scopes, 1982). In the case of a microbial source, one could try exposing the microbe to mutagens and hope that it would result in that particular microbe acquiring more desirable traits. Recombinant DNA technology changed the scene dramatically. Cloning for overexpression, site directed mutagenesis and directed evolution (Arnold and Georgiou, 2003) to obtain proteins with desired traits and specificity and tissue culture techniques (both for plant and animal tissues) have meant that sourcing of proteins can be done inside a laboratory! A more recent approach called metagenomics is further changing this upstream step in a significant way. This approach is based upon the realization that only a small percentage of microbes occurring in nature are cultivatable inside a laboratory. That means all these years, microbiologists have been missing out on much of the biodiversity! Again, using molecular biology, the environmental DNA can be isolated instead and used as a starting point to fish out proteins with desired characteristics (Roden et al., 1999; Torsvik and Goksoyr, 1980). This sudden abundance of sources for desirable proteins has necessitated a hard look at the way enzymologists were isolating proteins/enzymes. A widely quoted survey showed that most of the time enzymologists tend to follow a sequence of unit processes for protein purification. It consists of Precipitation → Ion exchange chromatography → Gel filtration → Affinity chromatography (Bonnerjea, 1986). Often, ion exchange chromatography is repeated a few times with different chromatographic media. An increasing

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number of steps in a purification protocol affects the ultimate yield of the pure protein (Mondal et al., 2006). Sadly, this practice continues and even (actually especially) hard core enzymology journals often publish such purification protocols. Such purification protocols are of no practical utility to biotechnologists who want to obtain such proteins in large amounts for bioconversions, bioanalysis or therapeutic applications (Przybycien et al., 2004; Smith, 2005). This is one area where lately many options have developed. It should be mentioned, at this stage, that for what is known as industrial enzymology/(or in recent years) white biotechnology, the purity of the enzyme need not be very high (Gupta and Raghava, 2007). For bioanalysis, the purity has to be a little higher. For therapeutic purposes, the protein has to be of very high purity (Table 2.1).

2.4.2

Chromatographic and nonchromatographic methods of protein purification

Before we review the current practices in large scale protein purification, it is useful to divide protein separation methods into two broad classes: chromatographic and nonchromatographic (Przybycien et al., 2004; Mondal et al., 2006; Mondal and Gupta, 2006). The chromatographic methods have high resolution and are costly. Table 2.2 lists a variety of chromatographic methods which are available today. It is not easy to scale up a chromatographic process. However, it is easier to put a chromatographic process onto a robotic platform (Smith, 2005) and this convenience associated with automation gets inbuilt in the production process. Nonchromatographic methods (listed Table 2.1 Purity of therapeutic proteins Impurity

Detection method(s)

Host cell and media proteins DNA Microbial load Turbidity Isoforms which have different properties than therapeutic protein Deamidation products Oxidation products Aggregated forms Proteolytic products Endotoxins

ELISA, SDS-PAGE Hybridization assay Bioburden testing Turbidometry Isoelectric focusing, SDS-PAGE, mass spectrometry Isoelectric focusing Tryptic mapping SDS-PAGE, Gel-filtration on HPLC SDS-PAGE, Gel-filtration on HPLC Limulus amebocyte lysate assay

The regulatory agencies in various countries require that the therapeutic protein should be checked for various contaminants (Harrison, 1994; Kelner and Bhagat, 2007)

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Natural-based polymers for biomedical applications Table 2.2 The variety of chromatographic methods Ion exchange chromatography Hydrophobic Interaction chromatography Reversed phase chromatography Hydroxyapatite chromatography Immobilized metal affinity chromatography Thiophilic interaction chromatography Hydrophobic charge induction chromatography Size exclusion chromatography Mixed mode ion exchange chromatography Affinity chromatography Perfusion chromatography Displacement chromatographya Simulated moving bed chromatographya Expanded bed adsorption chromatographya Chromatography on monoliths Membrane chromatographya a Notes: These are different modes of carrying out chromatography and can be carried out with most of the above listed versions provided appropriate chromatographic medium is available.

Table 2.3 Non chromatographic methods Non chromatographic methods Isoelectric precipitation Precipitation with salts, polymers, organic solvents, surfactants and metal ions Aqueous two phase separations Affinity precipitation Three phase partitioning Macro-(affinity ligand) facilitated three phase partitioning Crystallization Ultrafiltration

in Table 2.3) generally have low resolution, are easy to scale up and are generally less costly. Hence it follows that whenever a protein of very high purity is required, a chromatographic step would be invariably required. In other cases, it may be economical to use nonchromatographic methods.

2.4.3

Affinity based purification strategies

In the last 10 to 15 years, much effort has gone in reducing the number of unit processes used during protein purification. Most of it has relied upon the high selectivity of affinity ligands in binding to protein molecules (Gupta, 2002). During this period, the nature and designs of affinity ligands itself has

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been an area of hectic activity. Initially, it was found that a substrate analog/ coenzyme would bind the corresponding enzyme with high selectivity. Hence such an affinity ligand linked to an inert chromatographic material (generally a polysaccharide especially agarose) formed the basis for an affinity chromatography protocol. The column would bind the protein of choice selectively. Suitable methods for dissociation of the target protein from the affinity media quickly developed (Gupta, 2002). Change of pH to alter the binding constant or elution with a buffer containing the affinity ligand or its analog normally worked. Antibodies, naturally, were the obvious and possible affinity ligands as these could be generated against any given protein (Mondal and Gupta, 2006). At this stage of development, affinity chromatography was considered expensive because of the cost of the affinity ligand and the cost of the coupling reagents while creating the affinity medium. Hence, affinity chromatography was generally used at the last stage as a polishing step. At this stage the protein was in concentrated and nearly pure form and a small column of the affinity medium could be used. Two key developments changed this picture (Mondal and Gupta, 2006). Chronologically, these developments happened over a similar period and in a way helped each other. First, it was realized that affinity interactions could be interfaced with nonchromatographic methods. Hence, an affinity step (and its associated high selectivity) could be carried out very early in the protocol and even with dirty crude suspensions (as starting feed normally is). The chromatographic columns normally are prone to getting spoilt and choked with such crude feed/extracts (see, however, later discussion on expanded bed chromatography). Second was the realization that choice of an affinity ligand was not limited to those which have a biological (i.e. in vivo) relationship with the target protein (Mondal and Gupta, 2006). Thus availability of economical and robust affinity ligands (e.g. dyes and metal ions) and the possibility of obtaining in large amounts an affinity ligand (e.g. monoclonals, peptide libraries, aptamers) for any protein, expanded the range of affinity based separations in a significant way. Another major development was the approach of using affinity tags or the concept of fusion proteins (Przbycien, 2004; Mondal and Gupta, 2006; Mondal et al., 2006). In this approach, a part of DNA expressing an ‘affinity tag’ is fused with the gene expressing this desired protein/enzyme. This fusion protein binds to an appropriate affinity medium via the affinity tag. One of the most common affinity tags is the polyhistidine tag and protein containing polyhistidine tags is easily purified by passing through a chelated Ni column (see later for a description of this kind of chromatography which is called immobilized metal ion affinity chromatography (Roy et al., 2007)). Many of these technologies are commercially available as kits. When these work, they essentially reduce protein purification to a very simple exercise.

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Various techniques and their applications

Expanded bed chromatography A large number of chromatography media are available commercially. The choices available also provide the options of carrying out chromatography in low pressure, medium pressure (e.g. FPLC) and high pressure (HPLC) modes. Expanded bed chromatography is a mode which allows the direct processing of crude suspensions. (Roy et al., 2005; Sonnenfeld and Thömmes, 2007). It uses a fluidized bed but the hydrodynamic properties of the chromatography medium are such that it behaves like a packed bed (i.e. there is no back mixing of the beads). The suspended materials pass through the inter particle space and the protein(s) can be captured by the medium. The expanded bed chromatography can be carried out in the form of an ion exchange chromatography, hydrophobic interaction chromatography or even affinity chromatography. Unfortunately, the media marketed by Pharmacia Biotech tend to be very costly. Some cheaper options (Jain et al., 2002; Roy and Gupta, 2004; Roy et al., 2004) are described in the open literature but their use is limited as the current crop of biochemists are generally diffident about using anything which is not available in a ‘ready to use’ form. Another constraint with expended bed chromatography is identified in a concluding sentence of a recent article on the technique: ‘We encourage you to investigate EBA with an open mind and lots of math’ (Sonnenfeld and Thömmes, 2007). However, the advantages of expanded bed chromatography are worth the trouble as it is a purification technique that combines clarification (which requires filtration or centrifugation or a membrane filtration), concentration (which otherwise makes precipitation an obligatory initial step in protein purification) and chromatography. Immobilized metal ion affinity chromatography (IMAC) IMAC exploits the affinity of metal ions like Cu2+, Zn2+, Ni2+ for histidines, tryptophans, phenylalanines and tyrosines present on the protein surface (Kågedal, 1989; Deutscher, 1990). The positions of these residues and other protein surface properties also contribute to the retention behavior of the protein on an IMAC column. The metal ion is coordinated to a chelating ligand (two most commonly used ligands are iminodiacetic acid and nitrolotriacetic acid) which in turn is linked to one of the usual affinity supports (agarose being the most commonly used). While IMAC was originally developed to purify proteins exploiting the distribution of accessible histidine residues on the protein surface, its major application turned out to be purification of recombinant proteins with polyhistidine tag. Histidine residues constitute only about 2% of the amino acid contents of globular proteins. Furthermore, only 50% of these are exposed on the protein surface. Hence a recombinant

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protein with a polyhistidine tag can be very selectivity bound to an IMAC column (Mondal and Gupta, 2006). Aqueous two-phase separations (ATPS) While not generally used by enzymologists in the academic sector, ATPS has been widely used in industry for the last several decades (Johansson, 1989). It has been favoured because it is one of the very few techniques which can deal with crude suspensions directly without any preclarification. In ATPS, two immiscible phases are formed by mixing a polymer (e.g. PEG) with another polymer (e.g. dextran) or salt (e.g. K2SO4). PEG-salt systems are cheaper and hence preferred. In ATPS, a protein partitions between two phases depending upon its various properties such as charge, hydrophobicity, etc. The overall separation/recovery of a protein can be optimized by varying polymer, salt, ionic strength or pH (Johansson, 1989). It is possible to link an affinity ligand to PEG and incorporate PEG-affinity ligand conjugate into the PEG phase. This interface of an affinity interaction alters the partition coefficient of the target protein drastically (Mondal et al., 2006). An alternative approach is to place a conjugate of a smart polymer and the affinity ligand into the PEG phase (see below for a discussion on smart polymers in the affinity precipitation section) (Mondal et al., 2006). This allows easy separation of the protein from the PEG phase and reuse of this somewhat expensive polymer. Affinity precipitation This technique is based upon the fact that certain water soluble polymers called smart polymers or stimuli-responsive or stimuli-sensitive polymers (see Chapter 5 of this volume) can be conjugated to an appropriate affinity ligand. When mixed with a population of proteins, an affinity complex of smart polymer–affinity ligand with the desired protein is formed. Using the appropriate stimulus, this affinity complex can be precipitated out of the solution/suspension. The protein can be recovered from this affinity complex just as in affinity chromatography. Working with free solution has the advantage that the high selectivity of affinity interactions can be deployed right at the beginning. Being a precipitation technique, it also reduces the feed volumes and concentrate protein solutions. Hence, in principle, it can replace the ammonium sulfate precipitation which is a great favorite with traditional enzymologists. In practice, the two precipitation techniques can complement each other (Gupta, 2002; Mondal and Gupta, 2006; Mondal et al., 2006). In many cases, the smart polymers themselves have the inherent property of showing high selectivity towards the desired protein and no conjugation with an affinity ligand is required (Roy and Gupta, 2003; Mondal and Gupta,

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2006; Mondal et al., 2006). In such fortuitous cases, affinity precipitation constitute a very economical unit process. Macro-(affinity ligand) facilitated three phase partitioning (MLFTPP) Three phase partitioning (TPP) is another precipitation technique which exploits synergy of salt and an organic solvent to precipitate proteins (Sharma and Gupta, 2001a; Jain et al., 2004). The addition of ammonium sulphate and tbutanol to a protein create three phases: an upper t-butanol rich phase, a lower aqueous phase and an interfacial protein precipitate. TPP has been widely used for purification of a large number of proteins (Mondal and Gupta, 2006; Mondal et al., 2006). It was found that a smart polymer or its conjugate with an affinity ligand can also be floated as an interfacial precipitation just like proteins. This enables TPP to become an affinity based separation approach with high selectivity. When added to a mixture of protein during TPP, the smart bioconjugate floats between the two phases as a precipitate of its affinity complex with the target protein. The affinity complex can be dissociated and the desired protein recovered. Just like ATPS and expanded bed chromatography, MLFTPP can also be used directly with crude suspensions (Sharma and Gupta, 2002; Sharma et al., 2003). Membrane based separations Removal of low molecular substances from the protein solution/suspension by dialysis is the most familiar and the oldest example of the use of membranes in protein purification. Microfiltration operated with pressures of 0.1 – 1.0 bar retains particles in the range of 0.1 – 10 µm and is useful, for example in removing cell debris from the protein sample. Ultrafiltration operates with a higher pressure range of 0.5 – 10 bars and can separate proteins from each other (Lutz and Raghunath, 2007). Ultrafiltration (UF), also known as diafiltration (DF), is widely used to concentrate and exchange buffers of protein samples. The low molecular impurities are also, of course, removed. Again, linking affinity ligands to membranes has made, membrane separations more selective (Mondal et al., 2006). Various ultrafiltration modules: spiral type, hollow fiber type and cassette type, are available (Lutz and Raghunath, 2007; Mondal et al., 2006). These differ in performance characteristics such as robustness, mass transfer efficiency and scalability (all highest for cassette type). Hollow fiber type are however easiest to use and can work with very high feed flow rates. Modified polyether sulfone and regenerated cellulose are two commonly used membrane materials. Ultrafiltration is generally carried out in tangential flow filtration mode to minimize ‘fouling’ of the membrane which is caused by covering of the

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membrane layer with various retained molecules. Membrane chromatography has been developed to carry out affinity chromatography which allows increased flow rates without high pressure drops and high mass transfer efficiency. These days membrane chromatography is more frequently used for purification of plasmid DNA, viruses and very large proteins (>250 kDa) (Lutz and Raghunath, 2007). However, as pointed out by Przybycien et al. (2004), it has proved very useful in purification of therapeutic protein in removing trace impurities. In such cases, the therapeutic proteins are obtained in the flow through. The various applications of different techniques described above are listed in Table 2.4.

2.5

Purification of lipids

2.5.1

Extraction of lipids

There are some precautions which are necessary during extraction of the lipids irrespective of whether the source is an animal tissue, plant tissue or microbial in nature. Even before extraction, it is better to preserve the starting tissue at low temperature (< –20°C) and under inert atmosphere. The latter is especially desirable if the lipid material contains double bonds which are quite prone to oxidation. Hence for extracting lipids containing unsaturated fatty acids, it is advisable to add butylated hydroxytoluene (BHT) (1–10 mg/l) to the extracting solvent mixture. Enzymatic degradation is minimized by storage and processing at low temperature. Unfortunately it cannot be abolished. For example, enzymes in plant tissues are reported to cause degradation even at low temperatures (Deutscher, 1990). Exposure of the tissue to high temperature (e.g. boiling water or steam) can be tried to inactivate the problematic enzymes. Prior extraction with isopropanol also inactivates lipases and is recommended while working with plant tissues (Gurr and Harwood, 1991; Christie, 1990). Like other substances, the first step in obtaining lipids is extraction and generally requires mixture of solvents. An extraction with CHCl3: CH3OH: endogenous water (1:2:0.4) generally extracts all lipids. The tissue is homogenized in this one phase solvent system. Addition of any one of the component liquids leads to phase separation with nonlipids material partitioning in the upper aqueous phase. The lower CHCl3 phase can be washed with water and dried to obtain the lipid. Any residual water is removed by treating with anhydrous sodium sulphate. In the case of more polar lipids, water during extraction should be substituted with salt solution or dilute acid solution to ensure that such lipids do not partition into the upper phase. If only neutral lipids are required, the dried lipid sample can be extracted with cold dry acetone. Most of the phospholipid material is left behind (Christie, 1982, 1990).

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Table 2.4 Illustrative list of applications of various bioseparation strategies for proteins/enzymes Chromatographic methods

Proteins/enzymes

Reference

Gel-filtration

Hepatitis B surface antigen Luteinizing hormone

Belew et al. (1991) Chaudhary et al. (2006)

IMAC

Soybean trypsin inhibitor IgG

Gupta et al. (2002) Jain and Gupta (2004)

Ion exchange chromatography

IgG

Corthier et al. (1984)

Thiophilic interaction chromatography

IgA, IgG, IgM

Hutchens et al. (1990)

Affinity chromatography α-amylase Glycosylated hemoglobin

Sardar and Gupta (1998) Fluckiger et al. (1984)

Membrane chromatography

Lactoferrin

Ulber et al. (2001)

Monoliths

Clothing factor IX

Branovic et al. (2003)

Expanded bed chromatography

Alkaline phosphatase α-amylase α amylase/proteinase K inhibitor Polyphenol oxidase Pullulanase Jacalin

Roy and Gupta (2000) Roy et al. (2007) Roy and Gupta (2001)

Aqueous two phase separation

Xylanase and pullulanase Phospholipase D Chitinase Chitin binding lectins

Teotia Teotia Teotia Teotia

Affinity precipitation

Xylanase β-glucosidase Wheat germ α-amylase Phospholipase D Glucoamylase β-amylase Lipase Alcohol dehydrogenase

Gupta (1994) Agarwal and Gupta (1996) Sharma et al. (2000a) Sharma et al. (2000b) Teotia et al. (2001) Teotia et al. (2001a) Sharma and Gupta (2001d) Mondal et al. (2003c)

Roy et al. (2002) Roy and Gupta (2002) Roy et al. (2005) and Gupta (2001) and Gupta (2004) et al. (2004) et al. (2006)

Three phase partitioning Alkaline phosphatase Phospholipase D Protease/amylase inhibitor Pectinase Green fluorescent protein Xylanase

Sharma et al. (2000b) Sharma and Gupta (2001c) Sharma and Gupta (2001a) Sharma and Gupta (2001b) Jain et al. (2004) Roy et al. (2004)

MLFTPP

Xylanase Glucoamylase, pullulanase α-amylase

Sharma and Gupta (2002) Mondal et al. (2003b) Mondal et al. (2003a)

Membrane based methods

Lysozyme

Ghosh and Cui (2000)

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Fatty acids present in triglycerides constitute one of the industrially important products (Gupta, 1996). The mixture of fatty acids can be obtained by treating the tissue or lipid extract by dilute aqueous or methanolic KOH. Extraction with light petroleum gives nonsaponifiable lipids like sterols. Acidification of the extract with ether gives a mixture of fatty acids present in the tissue. The components of the lipid material at this stage will depend upon the nature of the tissue. Erythrocytes will contain phosphoglycerides, sphingolipids and sterols; animal liver will contain phosphoglycerides, sphingolipids, sterols and triglycerides; plant leaves will contain phospholipids, glycolipids, waxes and cutin; cyanobacteria will contain phosphoglycerides, glycolipids, waxes; gram negative bacteria will contain mostly phospholipids. Other solvent mixtures which have been tried are isopropanol: hexane (2:3) or chloroform: CH3OH in a different ratio (2:1). While extracting lipids from sources such as cereals wherein lipids form inclusion complexes with starch, butanol saturated with water is a useful solvent. This extraction is also useful for isolating lysophospholipids which are more soluble in water as compared to other phospholipids. Also, at this stage other impurities which are soluble in organic solvents are also expected to be present. Some of the nonlipid compounds which may be present in the extracted material are sugars, amino acids, urea and salts. The chloroform–ethanol extract shaken with one fourth of its volume of a 0.88% KCl solution (in water) forms two layers. The chloroform rich lower phase contains all the lipids material. The exception is gangliosides which go into the upper layer and can be recovered by dialyzing out low molecular weight impurities (Gunstone et al., 1986; Christie, 1990).

2.5.2

Further purification

Further purification of lipids requires adsorption column chromatography. Silica is frequently used as chromatographic media. The elution procedure for many lipids are now well worked out. Simple lipids can be eluted from silica gel with chloroform or diethylether, acetone elutes out glycolipids whereas methanol is used for eluting phospholipids. Acetone may also elute out phosphatidic acid, diphosphatidyl glycerol or even phosphatidylethanolamine. This can be avoided by incorporating CHCl3 in acetone. Methyl formate used before acetone elution would elute out prostaglandins with some amount of glycolipids. In practice, mixtures of solvents are tried to obtain optimum resolution with a given extract and a specific sample of column material. Other chromatographic media such as ion exchangers (for charged lipids) and boric acid bound to polymeric matrix (for glycolipids) have also been described in the literature (Christie, 1989). Among ion exchangers, DEAE-cellulose has been used most often at the preparative scale. The choline containing phospholipids get eluted with CHCl3: CH3OH mixture. Increasing CHCl3 ratio elutes phosphatidyl ethanolamine.

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Elution of phosphatidyl serine requires glacial acetic acid. These days a large number of ion exchangers are available in the market and ion exchangers are bound to be used more often for separation of lipids (Christie, 1990). Similarly complex separation of glycolipids is possible by using lectin columns. Lectins are proteins of nonimmune origin which show selective recognition of carbohydrate moieties. As a large number of lectins with different specificities are available, these constitute powerful tools for working with glycolipids (Van Damme et al., 1998).

2.5.3

Extraction and refining of edible fats/oils

Fats/oils used as a cooking media occupy an important role in human nutrition (Hui, 1996). Edible fats/oils are of either animal or plant origin. Many plant seeds constitute the source for the fat/oil. The cooking medium of choice tends to be different in various parts of the world. In big countries like India, even different part of the country tend to favor different oil/fat as a cooking medium. While specific industrial processes may vary depending upon the seed material or a particular industry, the following general discussion illustrates the various steps which are involved in obtaining fats/oils in ‘ready for cooking’ form (Anderson, 1996). After cleaning, most of the seeds have to be dehulled mechanically and may include the use of cracking rolls. In many cases, it is more practical to use hot dehulling wherein the moisture level of the seed is considerably reduced. After dehulling, flaking rolls are employed to break cell walls. After this, either mechanical pressing (for oil rich seeds such as sunflower or canola) or solvent extraction (for low oil content seeds such as soybeans) is carried out. In mechanical pressing, 60–90% oil is removed. Hexane is the solvent universally employed in solvent extraction of oils (Anderson, 1996). Oils rich in phosphorus (e.g. soybean, corn, sunflower) have to be degummed at this stage. Treatment of the oil with acids followed by hydration is effective in removing the gummy phospholipids. Enzymatic degumming with phospholipases can also be used (Andersson, 1996; Godfrey and West, 1996). Next treatment with alkali (‘caustic refining’) neutralizes free fatty acids to produce ‘soap’. This also helps in further degumming of oil. This is followed by bleaching (with clay) to produce refined oil suitable for marketing (Andersson, 1996). In view of solvent extraction with hexane producing ‘volatile organic compounds’ which constitute environmental hazards, many papers describing aqueous enzymatic oil extraction (AEOE) have been described (Sharma et al., 2001; Sharma et al., 2002b; Shah et al., 2005; Sharma and Gupta, 2006). In this approach, enzymes like proteases, cellulases, hemicellulases etc. have been used to liberate ‘oil bodies’ enmeshed in cellular structures. In another approach, TPP (discussed in Section 2.4) has also been used for extraction of

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oil from plant seeds (Sharma et al., 2002a; Shah et al., 2004; Sharma and Gupta, 2004; Gaur et al., 2007). For specific applications, vegetable oils, milk fat and animal fat (tallow and lard) is fractionated by multistage protocols which have been described in detail elsewhere (Krishnamurthy and Kellens, 1996). Finally, it may be added that lately there has been considerable interest in extracting nonedible oils from sources such as Jatropha for production of biodiesel (Francis et al., 2005; Shah et al., 2005; Shah and Gupta, 2007; Kumari et al., 2007).

2.6

Purification of polysaccharides

The first step here as well is the extraction of polysaccharides from the source material (Whistler, 1965). Some illustrative procedures are briefly described. The methods for extraction (and further purification) with particular references to some smart polysaccharides have been mentioned elsewhere in this book (see Chapter 5). Agar (agar-agar) is a galactoglycan and is obtained from species of the class Rhodophyceal (red purple seaweeds). Agar is soluble in hot water but insoluble in cold water. Hence, it is extracted by boiling water. After filtration/centrifugation, the solution is cooled to obtain the gel. This dissolution/cooling cycle is repeated a few times. Final washing with cold water and dehydration with absolute ethanol and acetone is followed by an ether wash to obtain agar. Acylation, fractional dissolution and deacylation gives agarose which is a pure and main component. Agarose is a linear molecule containing mostly galactose and 3,6 anhydro-L-galactose with small amount of sulfur and pyruvic acid. While agar is widely used in microbiology, agarose gel electrophoresis has become a key technique in molecular biology. Aqueous chloral hydrate has been reported to be an excellent solvent for extraction of bacterial starches. Precipitation with alcohol as a second step gives protein free starch of high purity. Chondroitin-4-sulfate is a polysaccharides consisting of alternating units of β-D-glucopyranosyluronic acid and 2-acetamido-2-deoxy -α-D-glucopyranosyl 4-sulfate units and is isolated from cartilage of skin, cornea, bone etc. Extraction with an alkaline solution, removal of proteins by treatment with kaolin, or phosphotungstic acid or amyl alcohol/CHCl3 and fractional precipitation with alcohol gives the product. A similar procedure can be employed for isolation of chondroitin6-sulfate from shark cartilage. Another class of industrially useful polysaccharides, dextrans, are isolated from cultures of leuconostoc mesenteroides and L-dextranicum. Both water soluble dextrans and water insoluble dextrans are known. Ethanol precipitation from the culture filtrates is the basic approach. Similar strategies are used for extraction of a large number of polysaccharides like galactans (from seeds, woods, seaweeds and as metabolic products of

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angiosperms), glycogen (from animals and some microoganisms), galactomannans (from guar seed), heparin (from animal tissues), hyaluronic acid (from connective tissues of animals), inulin (from tubers of dahlias or Jerusalem artichokes), levan (from bacteria) and xylans (from land plants and marine agar). In some cases, further fractionation when required can be carried out on ion exchangers, gel filtration columns, celite or calcium phosphate columns. Fractionation with quaternary ammonium salts or copper complexes has also been reported. Finally, one of most common polysaccharides, starch, as isolated, fractionated or derivatives is widely used in industry.

2.7

Purification of nucleic acids

The purification of nucleic acids is an integral step in most molecular biology experiments. Often, a small amount of DNA is isolated and amplified. The DNA preparations are used for cloning, southern blotting, PCR, real time PCR, random amplification of polymorphic DNA (RAPD), restriction fragment length polymorphism (RFLP) and amplified fragment length polymorphism (AFLP) analysis (Grossman and Moldave, 1968; Adams et al., 1986; Primrose et al., 2001).

2.7.1

Purification protocols used for isolation of DNA and RNA

One of the oldest techniques for isolation of DNA (Grossman and Moldave, 1968) is to mix cell lysate with phenol, CHCl3 and isoamyl alcohol. DNA partitions into the aqueous phase from where it is precipitated with alcohol. The purity is not high enough for applications such as PCR. Anion exchange chromatography is capable of yielding pure DNA. Such chromatographic media bind DNA selectively under low salt conditions as contaminants like metabolites, proteins and RNA are washed away. The DNA eluted with high salt buffers needs to be purified further by alcohol precipitation before high purity DNA is obtained. DNA of up to 150 kb can be purified by such protocols. DNA of up to 50 kb can be alternatively purified by selective binding to silica gel membrane in the presence of a high concentration of chaotropic salts. Low salt buffer eluted DNA is highly pure and does not even require an alcohol precipitation step. Based upon the above methods, a variety of kits are available from many vendors. RNA isolation Purification of intact RNA (Grossman and Moldave, 1968) is a key step in many molecular biology approaches like Northern analysis, RT-PCR, RNA mapping and cDNA library constructions. RNAses, as ubiquitous enzymes,

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pose the greatest threat to the integrity of RNA at all steps. To start with, cell disruption should be carried out in the presence of strong denaturants like LiCl, SDS or phenol which inactivate RNAses. All glassware/plasticware/ solutions need to be autoclaved. It is necessary to wear gloves to avoid nucleases from hands degrading the RNA. For isolation of total eukaryotic RNA, there are three common methods. In detergent/phenol extraction, the tissue is homogenized in the presence of detergent and the extract is mixed with either phenol and/or CHCl3. The nucleic acids partition in the aqueous phase for RNA. In guanidinium extraction, guanidinium thiocyanate + mercaptoethanol is used in the tissue homogenization buffer. In the proteinase K method, the enzyme is added to the tissue homogenization buffer containing SDS and this is followed by phenol/ CHCl3 treatment. In all three methods, ultracentrifugation employing a CsCl gradient is used to separate RNA as a pellet from DNA which has a lower buoyant density. Alternatively, pure DNAse can be used to hydrolyse away DNA. The total RNA obtained has to be separated when individual kinds of RNA are required (Sambrook and Russel, 2001). Eukaryotic mRNA (also called poly A RNA) constitutes about 1–5% of total cellular RNA and is required for synthesis of probes for array analysis and construction of random-primed cDNA libraries. The use of oligo (dT) as affinity ligand exploits the hybridization of oligo (dT) with poly A tails as with Watson and Crick’s well known base pairing. The prokaryotic mRNAs lack poly A tails and their separation from rRNA and tRNA is slightly tedious and often requires proprietary technology. tRNA on the other hand has a smaller molecular weight and can be separated by ion exchange chromatography. rRNA occurs as a part of ribosomes. Ribosomes can be separated as a cellular constituent by centrifugal of cell lysate. These, subjected to specific protocols, can give various rRNA which are constituents of ribosomes. It should be added that much RNA purification is carried out using specific kits available commercially. Kits tailored for a particular RNA and for specific applications are available. Also, robot based workstations are available for processing large numbers of samples for purification of a specific RNA.

2.7.2

Purification of plasmids

The major applications of plasmids are in cloning, gene therapy and production of DNA vaccines (Levi et al., 2000; Prather et al., 2003). As nonviral vectors, plasmids are less efficient in transfection. Thus, a large amount of plasmid DNA is required. Similarly, only one in 1000 plasmids reach the nucleus of the cell for expression of the therapeutic gene. For mass vaccination purposes, the plasmid DNA required is also enormous. Hence, unlike genomic DNA, large scale preparations of plasmids are required. Thus, while in principle,

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methods described for purification of DNA are also applicable for isolation of plasmid DNA, these have been generally replaced by large scale methods. A larger body of work deals with plasmid production from E. coli cells. Both batch and fed-batch technologies are available for plasmid overproduction by E. coli (Prather et al., 2003). It should be mentioned that very high purity plasmid DNA is required for gene therapy and vaccination. For example, the purity requirements for DNA vaccines demand that final plasmid DNA preparation should consist of > 90% closed plasmid DNA, < 1% genomic DNA, < 0.1% RNA, < 0.5 EU endotoxin and < 1% protein. (Prather et al., 2003). After cell lysis and separation of cell debris/other solids, a precipitation (by adding PEG, alcohols, cationic detergents or polyamines) step is generally followed by a chromatographic step. It is at this step that most of the protocols differ: anion exchanger, hydrophobic matrix, gel filtration media or an IMAC media (see discussion on IMAC in the section on protein purification) may be used. The plasmid eluted from such media is purified by precipitation. Let us discuss each step further. The method used for cell lysis plays an important role in the choice of the downstream steps which may be required. The most common technique is alkaline lysis with NaOH + SDS. The addition of sodium acetate precipitates proteins and genomic DNA. Unfortunately, any fragmented genomic DNA is not removed. Also, supercoiled plasmid DNA, in small amounts may be converted to denatured supercoiled, multimeric open and linear forms. An alternative protocol consisting of lysozyme treatment + boiling avoids such denaturation of plasmid DNA. The proteins are precipitated during boiling which also inactivates DNAses. Among precipitation approaches (which follow the cell lysis step), use of CTAB and spermidine is reported to give > 90% yields (Prather et al., 2003). The chromatographic step has been the subject of considerable research in recent years. Histidine – agarose, one of the affinity media described earlier for protein purification, has been used for selective purification of supercoiled plasmid DNA (Sousa et al., 2006). A patent which exploits triple helix chromatography in conjunction with chromatography on ceramic hydroxyapatite for obtaining pharmaceutical quality DNA has been described (Wils and Ollivier, 2004). As most of the available chromatographic media (for protein purification) have smaller pore sizes than plasmid DNA molecules, the capacity of such media tend to be only 0.2–2 g plasmid DNA per liter. Confocal microscopy showed that only 19% of the internal surface area of a protein purification chromatographic medium with 80 nm per diameter was available for plasmid purification (Danquah and Forde, 2007). Perfusion chromatography which utilizes media with ~ 4000 Å convective pores (which allow high flow rates) along with ~ 100 – 1000 Å smaller pores can be used with advantage (Levi et al., 2000). A customized biporous hydrophobic

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adsorbent for rapid purification of plasmid DNA with both high purity and yield has been described (Li et al., 2005). Tentacle chromatography wherein the media consists of long polyelectrolyte chains connected to the core matrix with long linkers is another promising approach (Prather et al., 2003). Monoliths (a continuous material consisting of organic or inorganic polymer with large pores) also constitute an interesting and alternative chromatographic media for plasmid purification. The suitability of a monolith ion exchanger for plasmid purification has been examined (Danquah and Forde, 2007). Among the nonchromatographic methods, aqueous two phase systems have been used in conjunction with either a membrane step (Frerix et al., 2006) or a chromatographic step (Trindade et al., 2005). A method for the direct purification of plasmid from yeast has also been described (Singh and Weil, 2002). The protocol is based upon a commercially available kit. In fact, such kits (in many cases) and many approaches described above can be used for purification of plasmid DNA from sources other than E. coli as well.

2.8

Purification of complex biomaterials

2.8.1

Miscellaneous complex biological molecules

Different classes of biological molecules (proteins, nucleic acids, lipids and carbohydrates) associate with each other to create cell organelles and necessary biological structure. Ribosomes, the sites of protein synthesis, are nucleoproteins and so are chromosomes. Biological membranes are basically lipoproteins; glycoproteins, proteoglycans and glycolipids are other important illustrative examples of complex biological molecules which all play important roles in living cells. ‘Soluble’ lipoproteins (in the blood serum) are important carriers of lipid molecules and are now well recognized as ‘good cholesterol’ (high density lipoproteins) and ‘bad cholesterol’ (low density lipoproteins). Classical methods of isolation of lipoproteins are fractional precipitation with ethanol/water mixtures, polyanions such as dextran sulfate or simple salts like ammonium sulfate. Ultracentrifugation with density gradients has emerged as a powerful method for fractionating lipoproteins (Gurr and Harwood, 1991). Preparation and characterization of plasma lipoproteins has been described in detail elsewhere (Segrest and Albers, 1986). Glycolipids can be eluted from the silica column with acetone, phospholipids requiring a more polar solvent with methanol (Gurr and Harwood, 1991). ATPS has been used for various membranes from mammalian tissue culture systems, as well as plant systems (Walter and Johansson, 1994). For purification of other complex biomolecules, readers should consult more specialized protocol books.

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2.8.2

Natural-based polymers for biomedical applications

Separation of animal cells

In a variety of contexts, it is necessary to separate different types of cells. While the methods mentioned here are, in principle, applicable to microbial cells and plant cells as well, the separation and fractionation of animal cells has attracted greater attention for obvious reasons. Separation of malignant cells from healthy cells and separation of different lymphocytes and their subclasses from each other are two illustrative examples. Cells are recognized by the collection of molecules present on their surfaces. Thus, the obvious way is to exploit the affinity of corresponding antibodies and lectins towards these surface molecules. The methods which exploit affinity infractions in cell separations have been reviewed by Hubble (1997). It is interesting to note that seminal work on ATPS in its title includes the term ‘cell particle’ (Albertsson, 1986). The protocol book on ATPS (Walter and Johansson, 1994) provide valuable protocols on separation and fractionation of cells by both simple partitioning and affinity partitioning. In fact, even the separation and subfractionation of organelles from both plant and animal cells are also discussed there. ATPS is a powerful tool to probe the surfaces of both prokaryotic and eukaryotic cells. The surface changes as a result of differentiation, maturation and aging can be probed. The cell-cell affinity can also be used for altering the partition coefficient of a particular cell type (Walter and Johansson, 1994). Hydrophobic or charged ligands have also been used. Chelated metal ions as ligands have been used to extend the concept of IMAC to affinity partitioning of cells (Walter and Johansson, 1994). Apart from ATPS, affinity ligands (antibodies, lectins, chelated metal ions, etc.) placed on magnetic beads have been used for fractionation of cells (Hubble, 1997). An exciting and emerging approach is to use smart materials for cell separation as well. Phase transitions in thermosensitive hydrogels used as graft on polystyrene dishes have been exploited for recovery of cultured cells (Yamada et al., 1990). Even physical coatings of smart polymers to polystyrene surfaces have been found to work well (Rollason et al., 1993). The conjugates of antibodies and thermosensitive polymers have been used in two phase affinity partitioning for separation of malignant cells (Gupta, 2002). The separation of animal cells continues to be a challenging area. Any development in this area is also going to impact on many other areas like tissue engineering and stem cell based applications.

2.9

Future trends

With the current thrust on sustainable development, naturally occurring biomaterials are becoming increasingly important. This increased importance is reflected in the coining of the new term white biotechnology. It has been

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estimated that 10–20% of all chemicals sold by the year 2010 will be made by biotechnological processes. So we are in the midst of a great paradigm shift. The trends in purification are a part of this big picture. It is obvious that increasingly, the starting point for many materials will be fermentation/ bioreactor. Biorefinery, a processing plant where renewable feedstock is converted into various valuable products, promises production of 30 building blocks which are key chemical intermediates in the chemical industry (Gupta and Raghava, 2007). Tapping renewable marine resources for a variety of polymers and low molecular weight materials is likely to be pursued vigorously. While the upstream part is going to witness remarkable changes, it is unlikely that we will see any new separation techniques. Biologists are increasingly working closely with biochemical engineers and the outcome would be further integration of the upstream and downstream component of the production process. A very good example which illustrates such an outcome is creating recombinant proteins with elastin like peptides (Roy et al., 2007). Such fusion proteins precipitate selectively on heating. Also, material science will play a greater role than ever before in designing more efficient adsorbents and chromatographic media. Down the line, even the biodegradability of separation media will become an issue. Sustainable practices will surely become important in separation science as well.

2.10

Acknowledgement

The preparation of this chapter and the research work from the author’s laboratory mentioned in this chapter were supported by Department of Science and Technology (Government of India) core group grant on ‘applied biocatalysis’ and Department of Biotechnology (Government of India) project grants.

2.11

Sources of further information

Barton D H R and Nakanishi K (eds) (1999), Comprehensive Natural Products Chemistry, Pergamon, Oxford, vol 1–9. Deutscher M P (ed.) (1990), Guide to Protein Purification, San Diego, Academic Press. Gupta M N (ed.) (2002), Methods for Affinity-based Separations of Enzymes and Proteins, Basel, Birkhauser Verlag. Hui Y H (ed.) (1996), Bailey’s Industrial Oil and Fat Products, New York, John Wiley, Vol 1–5. Sambrook J and Russel D W (ed.) (2001), Molecular Cloning: A Laboratory Manual, New York, Cold Spring Harbor Laboratory. Scope R K (ed.) (1982), Principles and Practice, Berlin, Springer-Verlag.

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Verrall M (ed.) (1996), Downstream Processing of Natural Products, Chichester, John Wiley. Whistler R L (ed.) (1965), Methods in Carbohydrate Chemistry, Vol V, New York, Academic Press.

2.12

References

Adams R L P, Knowler J T and Leader D P (1986), The Biochemistry of the Nucleic Acids, London, Chapman and Hall. Agarwal R and Gupta M N (1996), ‘Sequential precipitation with reversibly solubleinsoluble polymers as a bioseparation strategy: Purification of β-glucosidase from Trichoderma longibrachiatum’ Protein Expr Purif, 7, 294–298. Albertsson P A (1986), Partition of Cell Particles and Macromolecules, New York, Wiley. Alberts B, Bray D, Lewis J, Raft M, Roberts K and Watson J D (1994), Molecular Biology of the Cell, New York, Garland Publishing Inc. Alvarez-Lorenzo C and Concheiro A (2006), ‘Molecularly imprinted materials as advanced excipients for drug delivery system’, in El-Gewely M R (ed), Biotechnology Annual Review, Oxford, Elsevier, Vol. 12. Anderson L I, Nicholls I A and Mosbach K H (1994), ‘Molecular imprinting – a versatile technique for the preparation of separation materials of predetermined selectivity’, in Street G (ed), Highly Selective Separations in Biotechnology, Glasgow, Blackie Academic and Professional, 207–225. Andersson D (1996), ‘A Primer on oil processing technology’ in Hui Y H, Bailey’s Industrial Oil and Fat Products, New York, John Wiley, vol 4, 1–60. Arnold F H and Geogiou G (ed.) (2003), Directed Evolution Library Creation, New Jersey, Humana Press. Bailey J E and Ollis D F (1986), Biochemical Engineering Fundamentals, Singapore, McGraw-Hill. Belew M, Yafang M, Bin L, Berglof J and Janson J C (1991), ‘Purification of recombinant hepatitis B surface antigen produced by transformed Chinese hamster ovary (CHO) cell line grown in culture’, Bioseparation, 1, 397–408. Berg J M, Tymoczko J L and Streyer L (eds) (2002), Biochemistry, New York, W H Freeman. Best D J (1988), ‘Applications of biotechnology to chemical production in molecular biology and biotechnology’ in Walker J M and Gingold E, Molecular Biology and Biotechnology, London, Royal Society of Chemistry. Bonnerjea J, Oh S, Hoare M and Dunnill P (1986), ‘Protein purification: The right step at right time’, Bio/Technology, 4, 954–958. Branovic K, Buchacher A, Barut M, Strancar A and Josic D (2003), ‘Application of semiindustrial monolithic columns for downstream processing of clotting factor IX’, J Chromatogr B, 790, 175–182. Chaudhary R, Jain S, Muralidhar K and Gupta M N (2006), ‘Purification of bubaline luteinizing hormone by gel filtration chromatography in the presence of Blue Dextran’ Process Biochem, 41, 562–566. Chen X, Lin Y, Liu M and Gilson M K (2002), ‘The binding database: data management and interface design’, Bioinformatics, 18, 130–139. Christie W W (1982), Lipids analysis, Oxford, Pergamon Books.

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Christie WW (ed.) (1989), Gas Chromatography and Lipids: A Practical Guide, Somerset, The Oily Press. Christie W W (1990), Gas Chromatography and Lipids, Somerset, The Oily Press. Cortheir G, Boschetti E and Charley-Poulain J (1984), ‘Improved method for IgG purification from various animal species by ion exchange chromatography’, J Immunol Meth, 66, 75–79. Danquah M K and Forde G M (2007), ‘The suitability of DEAE-Cl active groups of customized poly (GMA-co-EDMA) continuous stationary phase for fast enzyme-free isolation of plasmid DNA’, J Chromatogr B, doi: 10.1016/j.jchromb. 2007.02.050. Davidson J N (ed.) (1965), The Biochemistry of the Nucleic Acids, London, English Language Book Society. Deutscher M P (ed.) (1990), Guide to Protein Purification, Methods in Enzymology, New York, Academic Press Inc, vol. 182. Feil H, Bae Y H, Feijen J and Kim S W (1991), ‘Molecular separation by thermosensitive hydrogel membranes’, J Membr Sci, 64, 283–294. Finar I L (1964), Organic Chemistry, London, Longmans, Green and Co Ltd., Vol 2. Fleming H L (1992), ‘Consider membrane pervaporation’, Chem Eng Prog, 88, 46–52. Fluckiger R, Woodtli T and Berger W (1984), ‘Quantitation of glycosylated hemoglobin by boronate affinity chromatography’, Diabetes, 33, 73–76. Francis G, Edinger R and Becker K (2005), ‘A concept for simultaneous wasteland reclamation, fuel production and socio-economic development in degraded areas in India: Need potential and perspectives of Jatropha plantations’, Nat Resour Forum, 29, 12–24. Frerix A, Geilenkirchen P, Müller M, Kula M R and Hubbuch J (2006), ‘Separation of Genomic DNA, RNA and Open Circular Plasmid DNA from Supercoiled plasmid DNA by combining denaturation, selective Renaturation and Aqueous Two-Phase Extraction’, Biotechnol Bioeng, 96, 57–66. Frost R (1999), ‘Enzyme Model’, in Dugas H (ed.), Bioorganic Chemistry: A Chemical Approach to Enzyme Action, New York, Springer-Verlag, 252–387. Furth A J (ed.) (1980), Lipids and Polysaccharides in biology, London, Arnold. Gaur R, Sharma A, Khare S K and Gupta M N (2007), ‘A novel process for extraction of edible oils. Enzyme assisted three phase partitioning (EATPP)’, Bioresour Technol, 98, 696–699. Ghosh R and Cui Z F (2000), ‘Purification of lysozyme using ultrafiltration’, Biotechnol Bioeng, 68, 191–203. Glazer A N and Nikaido H (1995), Microbial Biotechnology, New York, W H Freeman. Godfrey T and West S (eds) (1996), Industrial Enzymology, New York, Stockton Press. Grossman L and Moldave K (eds) (1968), Nucleic acids. (Methods in Enzymology, Vol XII Part B), New York, Academic Press. Gunstone F D, Harwood J L and Padley F B (ed.) (1986), The Lipid Handbook, London, Chapman and Hall. Gupta M N, Guoqiang D, Kaul R and Mattiasson B (1994), ‘Purification of xylanase from Trichoderma viride by precipitation with an anionic polymer Eudragit S-100’ Biotechnol Techniq, 8, 117–122. Gupta M (1996), ‘Manufacturing process for emulsifiers’ in Hui Y H (ed.) Bailey’s Industrial Oil and Fat Products, New York, John Wiley, vol. 4, 569–601. Gupta M N (ed.) (2002), Methods in Affinity-based Separation of Proteins/enzymes, Basel, Birkhauser Verlag.

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Gupta M N, Jain S and Roy I (2002), ‘Immobilized metal affinity chromatography without chelating ligands – Purification of soybean trypsin inhibitor on zinc alginate beads’, Biotechnol Prog, 18, 78–81. Gupta M N and Raghava S (2007), ‘Relevance of chemistry to white biotechnology’, Chem Central J, 1: 17. Gurr M I and Harwood J L (1991), Lipid Biochemistry. An Introduction, London, Chapman and Hall. Harrison R G (ed.) (1994), Protein Purification Process Engineering, New York, Marcel Dekker. Heftmann E (ed.) (1974), Chromatography, New York, Reinhold. Heritage J, Evans E G V and Killington R A (1996), Introductory Microbiology, Cambridge University Press. Hirotsu T (1987), ‘Water–ethanol separation by pervaporation through plasma graft polymerized membrane’, J Appl Polym Sci, 34, 1159–1172. Hostettmann K and Marston A (1997), Preparative Chromatography Techniques: Applications in Natural Products Isolation, Heidelberg, Springer. Hubble J (1997), ‘Affinity cell separations: problems and prospects’, TIBTECH, 15, 249– 255. Hui Y H (ed.) (1996), Bailey’s Industrial Oil and Fat Products, New York, John Wiley, Vol. 4. Hutchens T W, Magnuson J S and Yip T T (1990), ‘Secretory IgA, IgG and IgM immunoglobulins isolated simultaneously from cloistral whey by selective thiophilic adsorption’, J Immunol Meth, 128, 89–99. Jain S, Singh R and Gupta M N (2004), ‘Purification of recombinant green fluorescent protein by three phase partitioning’, J Chromatogr A, 1035, 83–86. Jain S and Gupta M N (2004), ‘Purification of goat IgG by immobilized metal ion affinity using crosslinked alginate beads’, Biotechnol Appl Biochem, 39, 319–322. James A T and Morris L J (eds) (1964), New Biochemical Separations, London, D. van Nostrand. Jenck J F, Agterberg F and Droescher M J (2004), ‘Products and processes for a sustainable chemical industry: a review of achievements and prospects’, Green Chem, 6, 544– 556. Johansson G (1989), ‘Affinity partitioning of proteins using aqueous two-phase systems’ in Janson J C and Ryden L (eds), Protein Purification: Principles, High Resolution Methods and Applications, Sweden, VCH Publishers, 330–345. Johnson A R and Davenport J B (eds) (1971), Biochemistry and Methodology of Lipids, New York, John Wiley. Kågedal, L (1989), ‘Immobilization meta ion affinity chromatography’ in Janson J C and Ryden L (eds), Protein Purification: Principles, High Resolution Methods and Applications, Sweden, VCH Publishers, 227–251. Kelner D N and Bhagat M K (2007), ‘Analytical strategy for biopharmaceutical development’, in Shukla A A; Etzel M R and Gadam S, (eds) Process Scale Bioseparation for the Biopharmaceutical Industry, Boca Raton, CRC Press, 395–418. Kim J J and Park K (1999), ‘Smart hydrogels for bioseparation’, Bioseparation, 7, 177– 184. Krishnamurthy R and Kellens M (1996), ‘Fractionation and Winterization’ in Hui Y H (ed.), Bailey’s Industrial Oil and Fat Products, New York, John Wiley, Vol. 4, 301– 337.

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Kumari V, Shah S, Gupta M N (2007), ‘Preparation of biodiesel by lipase catalyzed transesterification of high free fatty acid containing oil from Madhuca indica’, Energy Fuels, 21, 368 –372. Levy M S, O’Kennedy R D, Ayazi-Shamlou P and Dunnill P (2000) ‘Biochemical engineering approaches to the challenges of producing pure plasmid DNA’, TIBTECH, 18, 296–305. Li Y, Dong X Y and Sun Y (2005), ‘High-speed chromatographic purification of plasmid DNA with a customized biporous hydrophobic adsorbent’, Biochem Eng J., 27, 33– 39. Lutz H and Raghunath B (2007), ‘Ultrafiltration process design and implementation’ in Shukla A A; Etzel M R and Gadam S, (eds), Process Scale Bioseparation for the Biopharmaceutical industry, Boca Raton, FL, CRC Press, 297–332. Mondal K, Mehta P and Gupta M N (2003d), ‘Affinity precipitation of Aspergillus niger pectinase by microwave treated alginate’, Protein Expr Purif, 33, 104–109. Mondal K, Roy I and Gupta M N (2003c), ‘κ-Carrageenan as a carrier in affinity precipitation of yeast alcohol dehydrogenase’, Protein Expr Purif, 32(1), 151–160. Mondal K, Sharma A and Gupta M N (2003b), ‘Macro-(affinity ligand) facilitated three phase partitioning (MLFTPP) for purification of glucoamylase and pullulanase using alginate’ Protein Expr and Purif, 28(1), 190–195. Mondal K, Sharma A and Gupta M N (2003a), ‘Macro-(affinity ligand) facilitated three phase partitioning (MLFTPP) of α-amylases using modified alginate’ Biotechnol Prog, 19, 493–494. Mondal K, Gupta M N (2006), ‘The affinity concept in bioseparation: Evolving paradigm and expanding range of applications’, Biomol Eng, 23, 59–76. Mondal K, Roy I and Gupta M N (2006), ‘Affinity based strategies for protein purification’, Anal Chem, 78, 3499–3504. Nonaka T, Ogata T and Kurihara S (1994), ‘Preparation of poly (vinyl alcohol)-graft – Nisopropylacrylamide copolymer membranes and permeation of solvents through the membrane’, J Appl Polym Sci, 52, 951–957. Prather K J, Sagar S, Murphy J and Chartrain M (2003), ‘Industrial scale production of plasmid DNA for vaccine and gene therapy: plasmid design, production and purification’ Enzyme Microb Technol, 33, 865–883. Primrose S B, Twyman R M and Old R W (eds) (2001), Principles of Gene Manipulation, Oxford, Blackwell Publishing Company. Przybycien T M, Pujar N S and Steele L M (2004), ‘Alternative bioseparation operations: life beyond packed-bed chromatography’, Curr Opin Biotechnol, 15, 469–478. Roden M R, Goodman R M and Handelman J (1999), ‘The earth’s bounty: accessing soil microbial diversity, TIBTECH, 17, 403–409. Rollason G, Davies J E and Sefton M V (1993), ‘Preliminary report on cell culture on a thermally reversible copolymer’, Biomaterials, 14, 153–155. Roy I and Gupta M N (2000), ‘Purification of alkaline phosphatase from chicken intestine by expanded bed affinity chromatography on dye-linked cellulose’, Biotechnol Appl Biochem, 32, 81–87. Roy I, Sastry M S R, Johri B N and Gupta M N (2001), ‘Purification of alpha amylase isoenzymes from Scytalidium thermophilum on a fluidized bed of alginate beads followed by Concanavalin A-agarose column’, Protein Expr Purif, 20, 162–168. Roy I and Gupta M N (2001), ‘Purification of a ‘double-headed’ inhibitor of α-amylase/ Proteinase K from wheat germ by expanded bed chromatography’, Bioseparation, 9, 239–245.

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Roy I, Sharma S and Gupta M N (2002), ‘Separation of an isoenzyme of polyphenol oxidase from Duranta plumieri by expanded bed chromatography’, Protein Expr Purif, 24(2), 181–187. Roy I and Gupta M N (2002), ‘Purification of a bacterial pullulanase on a fluidized bed of calcium alginate beads’, J Chromatogr A, 950(1–2), 131–137. Roy I and Gupta M N (2003), ‘Smart polymeric materials: emerging biochemical applications’, Chem Biol, 10, 1161–1171. Roy I and Gupta M N (2004), ‘Hydrolysis of starch by a mixture of glucoamylase and pullulanase entrapped individually in calcium alginate beads’ Enzyme Microb Technol, 34, 26–32. Roy I, Jain S, Teotia S and Gupta M N (2004), ‘Evaluation of micro beads of calcium alginate for affinity chromatography of Aspergillus niger pectinase’, Biotechnol Prog, 20, 1490–1495. Roy I, Sardar M and Gupta M N (2005), ‘Crosslinked alginate-guar gum beads as fluidized bed affinity media for purification of jacalin’, Biochem Eng J, 23, 193–198. Roy I, Mondal K and Gupta M N (2007), ‘Leveraging protein purification strategies in proteomics’, J Chromatogr B, 849, 32–42. Ruckenstein E and Sun F (1995), ‘Concentrated emulsion pathway to novel composite membranes and their use in pervaporation’, Ind Eng Chem Res, 34, 3581–3589. Sambrook J and Russel D W (ed.) (2001), Molecular Cloning: A Laboratory Manual, New York, Cold Spring Harbor Laboratory. Sardar M and Gupta M N (1998), ‘Alginate beads as an affinity material for α-amylases’, Bioseparation, 7, 159–165. Scope R K (ed.) (1982), Protein Purification, Principle and Practice, New York, SpringerVerlag. Segrest J P and Albers J J (eds) (1986), ‘Plasma lipoprotein (Part A: preparation, structure and molecular biology) Method in Enzymology, New York, Academic Press, vol 128. Shah S, Sharma S and Gupta M N (2003), ‘Enzymatic transesterification for biodiesel production’, Indian J Biochem. Biophys, 40, 393–399. Shah S, Sharma S and Gupta M N (2004a), ‘Biodiesel preparation by lipase catalyzed transesterification of Jatropha oil’, Energy Fuels, 40, 1077–1082. Shah S, Sharma A and Gupta M N (2004b), ‘Extraction of oil from Jatropha curcas L seed kernels by enzyme assisted three phase partitioning’, Ind Crop Prod, 20, 275– 279. Shah S, Sharma A and Gupta M N (2005), ‘Extraction of oil from Jatropha curcas L seed Kernels by combination of ultrasonication and aqueous enzymatic oil extraction’, Bioresour Technol, 96, 121–123. Shah S and Gupta M N (2007), ‘Lipase catalyzed preparation of biodiesel from Jatropha oil in a solvent free system’, Process Biochem, 42, 409–414. Sharma A, Sharma S and Gupta M N (2000a), ‘Purification of wheat germ amylase by precipitation’, Protein Expr Purif, 18, 111–114. Sharma A, Sharma S and Gupta M N (2000b), ‘Purification of alkaline phosphatase from chicken intestine by three-phase partitioning and use of Phenyl-Sepharose 6B in the batch mode’ Bioseparation, 9, 155–161. Sharma A and Gupta M N (2001a), ‘Three phase partitioning as a large-scale separation method for purification of a wheat germ bifunctional protease/amylase inhibitor’, Process Biochem, 37, 193–196. Sharma A and Gupta M N (2001b), ‘Purification of pectinase from tomato using three phase partitioning’, Biotechnol Lett, 23, 1625–1627.

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Sharma A, Khare S K and Gupta M N (2001), ‘Enzyme-assisted aqueous extraction of oil from peanut seeds’, J Am Oil Chem Soc, 78, 949–951. Sharma A and Gupta M N (2002), ‘Macro-(affinity ligand) facilitated three phase partitioning (MLFTPP) for purification of xylanase’, Biotechnol Bioeng, 80, 228–232. Sharma A, Khare S K and Gupta M N (2002a), ‘Three phase partitioning of extraction of oil from soybean’, Bioresour Technol, 85, 327–329. Sharma A, Khare S K and Gupta M N (2002b), ‘Enzyme-assisted aqueous extraction of rice bran oil’, J Am Oil Chem Soc, 79, 215–218. Sharma A, Mondal K and Gupta M N (2003), ‘Separation of enzymes by sequential macro-(affinity ligand) facilitated three phase partitioning’ J Chromatogr A, 995, 127–134. Sharma A, Roy I and Gupta M N (2004), ‘Affinity precipitation and macro-(affinity ligand) facilitated three phase partitioning for refolding and simultaneous purification of urea-denaturated pectinase’, Biotechnol Prog, 20, 1255–1256. Sharma A and Gupta M N (2004), ‘Extraction oil from almond, apricot and rice bran by three phase partitioning after ultrasonication’, Eur J Lipid Sci Technol, 106, 183–186. Sharma A and Gupta M N (2006), ‘Ultrasonic pre-irradiation effect upon aqueous enzymatic oil extraction from almond and apricot seeds’, Ultrason Sonochem, 13, 529–534. Sharma S, Sharma A and Gupta M N (2000), One step purification of peanut phospholipase D by precipitation with alginate, Bioseparation, 9, 93–98. Sharma S and Gupta M N (2001c), ‘Purification of phospholipase D from Dacus carota by three-phase partitioning and its characterization’ Protein Expr Purif, 21, 310–316. Sharma S and Gupta M N (2001d), ‘Alginate as a macro-(affinity ligand) and an additive for enhanced activity and thermostability of lipases’, Biotechnol Appl Biochem, 35, 161–165. Shukla A A and Yigzaw Y (2007), ‘Modes of preparative chromatography’ in Shukla A A; Etzel M R and Gadam S, (eds), Process Scale Bioseparation for the Biopharmaceutical Industry, Boca Raton, FL, CRC Press, 179–220. Singh M V, Weil P A (2002), ‘A method for plasmid purification directly from yeast’ Anal Biochem, 307, 13–17. Smith C (2005), Striving for purity: advances in protein purification, Nat Methods, 2, 71– 76. Sonnenefeld A and Thömmes J (2007), ‘Expanded bed adsorption for capture from crude solution’ in Shukla A A; Etzel M R and Gadam S, (eds) Process Scale Bioseparation for the Biopharmaceutical Industry, Boca Raton, FL, CRC Press, 59–81. Sousa F, Freitas S, Azzoni A R, Prazeres D M F and Queiroz J (2006), ‘Selective purification of supercoiled plasmid DNA from clarified cell lysates with a single histidine-agarose chromatography step’, Biotechnol Appl Biochem, 45, 131–140. Subramanian G (ed.) (1998), Bioseparation and Bioprocessing A Handbook, Weinheim, Wiley VCH, vol I. Teotia S and Gupta M N (2001), ‘Free polymeric bioligands in aqueous two phase affinity extractions of microbial xylanases and pullulanase’, Protein Expr Purif, 22, 484– 488. Teotia S, Khare S K and Gupta M N (2001a), ‘An efficient purification process for sweet potato β-amylase by affinity precipitation with alginate’ Enzyme Microb Technol 28, 792–795. Teotia S, Lata R, Khare S K and Gupta M N (2001b), ‘One step purification of glucoamylase by affinity precipitation with alginate’ J Mol Recogn, 14, 295–299. Teotia S, Lata R and Gupta M N (2004), ‘Chitosan as a macro-(affinity ligand) purification

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of chitanases by affinity precipitation and aqueous two phase extraction’, J Chromatogr A, 1052, 85–91. Teotia S and Gupta M N (2004), ‘ Purification of phospholipase D by two-phase affinity extraction’, J Chromatogr A, 1025, 297–301. Teotia S, Mondal K and Gupta M N (2006), ‘Integration of affinity precipitation with partitioning methods for bioseparation of chitin binding lectins’, Food Bioprod Process, 84, 37–43. Thornton J D (ed.) (1992), Science and Practice of Liquid-liquid Extraction, Oxford, Oxford University Press. Torsvik V L and Goksoyr J (1980), ‘Determination of bacterial DNA in soil’, Soil Biol Biochem, 10, 7–12. Trindade I P, Diogo M M, Prazeres D M F and Marcos J C (2005), ‘Purification of plasmid DNA vectors by aqueous two-phase extraction and hydrophobic interaction chromatography’ J Chromatogr A, 1082, 176–184. Ulber R, Plate K, Weiss T, Demmer W, Buchholz H and Scheper T (2001), ‘Downstream processing of bovine lactoferrin from sweet whey’, Acta Biotechnol, 21, 27–34. Van Damme E J M, Peumans W J, Pusztai A and Bardocz S (eds) (1998), Handbook of Plant Lectins: Properties and Biomedical Applications, Chichester, John Wiley and Sons. Verrall M (ed.) (1996), Downstream Processing of Natural Products, Chichester, John Wiley. Walter H and Johansson G (ed.) (1994), Aqueous Two Phase Systems, New York, Academic Press. Weissberger A (ed.) (1965), Techniques of Organic Chemistry, New York, Interscience, Vols IV–VI. Whistler R L (ed.) (1965), Methods in Carbohydrate Chemistry, New York Academic Press, Vol V. Wils P and Ollivier M (2004), ‘Purification of plasmid DNA of pharmaceutical quality’, US patent 6730781 B1. Yamada N, Okano T, Sakai H, Karikusa F Sawasaki Y and Sakurai Y (1990), ‘Thermoresponsive polymeric surfaces; control of attachment and detachment of cultured cells,’, Macromol Chem Rapid Commun, 11, 571–576.

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3 Processing of starch-based blends for biomedical applications R. A. S O U S A, V. M. C O R R E L O, S. C H U N G, N. M. N E V E S, J. F. M A N O and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

3.1

Introduction

Nature has produced a myriad of polymers with large biomedical potential. Natural polymers can be broadly categorized in eight different categories, namely: (1) polysaccharides, (2) proteins, (3) polyhydroxyalkanoates, (4) polythioesters, (5) polyanhydrides, (6) polyisoprenoids, (7) lignin and (8) nucleic acids. Many biodegradable formulations based on these natural polymers have been developed. One of the main disadvantages of biodegradable polymers obtained from natural derived polymers is their predominant hydrophilic nature, which results in inherent fast degradation rates, but, in many cases, poor mechanical performance. These properties can be significantly improved, in many cases, by blending the natural polymers with other biodegradable polymers from synthetic origin. So far, our Research Group has developed significant work on what concerns the processing and characterization of polysaccharide based materials for biomedical applications, with special emphasis to the processing of blends based on corn starch. In this chapter, the processing and characterization of several blends based on starch is described and discussed in the context of several potential applications within the biomedical field.

3.2

Starch

Starch is the main carbohydrate reserve of higher plants, where it is found in storage organs such as seeds and tubers.1 Starch comes mostly from a small number of crops, namely maize, potato, wheat and tapioca, as well as from rice, sorghum, sweet potato, arrowroot, sago, and mung beans.1 Starch can be fractionated into two types of macromolecules: amylose and amylopectin. Amylose is itself a linear molecule of (1→4) linked α-D-glucopyranosyl units, slightly branched by (1→6)-α-linkages, while amylopectin is a highly branched molecule, containing both (1→4)-α-linkages bonds and (1→6)-αlinkages, at 25–30 glucose units distance. Amylose has a molecular weight 85 © 2008, Woodhead Publishing Limited

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of approximately 1 × 105–1 × 106 while amylopectin is a larger molecule, with molecular weight in the range 1 × 107–1 × 109.2–8 For most cereal starches, the relative weight percentages of amylose and amylopectin varies between 72 and 82% amylopectin, and 18 and 33% amylose.7 However, mutations can significantly affect the amount of both molecules in starch.1,7 Due to the different molecular structures and molecular weight ranges of amylose and amylopectin, starches with dissimilar amylose/amylopectin content ratios can exhibit significant differences in terms of properties.6,9–11 Starch is semi-crystalline, exhibiting a degree of crystallinity between 15% and 45%.12 Starch structure can be classified to A, B and C forms. In the native form, the A pattern is predominantly associated with cereal starches, while the B form is usually obtained from tuber starches. The A form adopts a close-packed arrangement with water molecules between each double helical structure, while the B-type is more open with more water molecules, essentially all of which are located in a central cavity surrounded by six double helices. The C form is a mixture of both A and B types that can be found in bean starches.13–16 Starches contain also phospholipids and free fatty acids which are positively correlated with the amylose fraction. Complexation of amylose by aliphatic fatty acids, emulsifiers or iodine result in V-type conformation.2,13

3.2.1

Gelatinization of starch

The original application of starch was found in the food industry.17 However, starch has also been recognized as a potential functional raw material in many other applications,17–19 including biomedical (which is covered later in this chapter). In order to meet varied applications, starch must be adequately modified by destructuring its granular structure. When heated in the presence of water, starch undergoes an irreversible order–disorder transition designated as gelatinization,20 in which the granules are observed to swell, absorb water, lose crystallinity, and to leach amylose. Gelatinization temperature is reached where these destabilized crystalline regions melt, leading to an irreversible loss of the granule structure.21,22 Gelatinization ultimately results in the formation of a viscous paste with disruption of most inter-molecular hydrogen bonds.23–26 The plasticization of starch is accomplished upon the fragmentation of the crystalline structure within the polysaccharides, by converting native starch granules to a highly amorphous paste, which is able to be processed as a thermoplastic formulation by conventional extrusion or injection moulding methods.23,27–29 As the glass transition temperature (Tg) and the melting temperature of pure dry starch are higher than its decomposition temperature, starch plasticization requires a plasticizer aimed at ensuring that starch undergoes gelatinization instead of degradation.30–33 Thermoplastic starch (TPS) can be obtained by an adequate combination of high pressure, high

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temperature and high shear conditions in the presence of water and/or other plasticizers.12,17,23,27,29 In this context, water is a commonly used plasticizer for the processing of starch. Addition of other plasticizers such as glycols, sugars and amides can also lower the Tg, by spacing out the molecules and reducing the intermolecular interactions. Distinct factors determine the final properties of TPS products, such as: molecular weight, the amylose/amylopectin ratio, the crystallinity of the products, and the type and amount of plasticizers.34–37 De Graaf et al.34 studied the interrelated effect of the plasticizer and the amylose/ amylopectin ratio into the mechanical properties of four different starches (potato, pea, wheat and waxy maize). The results indicate that Tg diminishes upon increase in glycerol content. In terms of mechanical properties, the modulus and tensile strength also decrease while the elongation enhances for higher glycerol contents. In terms of the amylose/amylopectin ratio, an increase in this factor lowers the modulus and tensile strength and increases ductility. Hulleman et al.35 studied the mechanical performance of different compression moulded mixtures of starches (corn, potato, waxy corn and wheat starch) and glycerol. The mechanical performance was strongly dependent on the water content of the premixes and on the starch source. Van Soest et al.36 investigated the structure and the mechanical properties of compression-moulded normal and high-amylose content maize starches as a function of processing water content and ageing time. The materials from high amylase maize starches are tougher and exhibit higher strength and lower elongations as compared to the normal maize starch materials. Differences in mechanical performance were attributed to differences in amylose content and to differences in branching of amyloptecin. Lourdin et al.37 studied the sorption behaviour and glass transition of starch films plasticized with varying concentrations of different plasticizers (glycerol, sorbitol, lactic acid sodium, urea, ethylene glycol, diethylene glycol, polyethylene glycol and glycerol diacetate). Glass transition generally decreased upon plasticizer increase. However, the plasticization effect was suggested to be dependent on favourable interactions (hydrogen bonds) with starch.

3.2.2

The thermosensitive character of starch

Starch-based materials are highly thermosensitive materials which easily degrade at high shear rates and with long residence times during processing. The processing of starch by extrusion or injection moulding produces lower molecular weight polysaccharides due to shear induced fragmentation. Molecular weight degradation has been shown to increase with increasing specific mechanical energy (SME) during processing.38–43 In this regard, Sagar, et al.44 reported the susceptibility of starch to thermo-mechanical degradation, by showing that thermo-mechanical conditions imposed during extrusion can cause macromolecular debranching and consequent decrease

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of molecular weight of high amylose starch. Vergnes, et al.45 reported the same effect for corn starch, upon studying the respective rheological behaviour using a pre-shearing rheometer. In this rheometer, it was possible to induce precise thermo-mechanical environments to the melt before viscosity measurements were taken. Starch degradation, characterized by the water solubility and intrinsic viscosity measurements, was shown to depend on the mechanical energy involved in the thermo-mechanical processing operations. Cunningham 46 reported the effect of several factors, such as starch concentration, temperature and screw speed on the intrinsic viscosity of extruded corn starch. Extrusion temperature proved to have the strongest effect on the starch extrudate viscosity, determining the level of thermomechanical degradation during processing. Intermediate temperatures were found to be a compromise between the excessive shear heating at low temperatures (which causes mechanical degradation) and the excessive thermal degradation at high temperatures. The occurrence of molecular degradation in starch during extrusion has been correlated with the specific mechanical energy (SME) input during melt processing.41–43 In this matter, Willet et al.41 studied the melt rheology and the degree of degradation of waxy maize starch following two consecutive extrusion passes after the conversion from native granules. The molecular weight of starch was found to decrease with increasing SME input, in agreement with other investigations,42,43 which have concluded that molecular weight decreases upon fragmentation of starch during extrusion. This is more pronounced for amylopectin molecules with higher molecular weight, as these are more prone to fragmentation/debranching during shear. The molecular weight decrease was also observed for starch in wheat flours during twin-screw extrusion, where a significant inverse relationship between SME and molecular weight data was found to occur.43 In another study, Brümmer et al.47 investigated the influence of the SME on the molecular structure of extruded starch. The chromatographic examination of the molecular changes in the starch revealed that SME had a significant positive effect on molecular degradation for lower processing temperatures (110–180°C).

3.3

Starch-based blends

The application of TPS is limited by the inherent susceptibility to thermomechanical degradation. In order to circumvent this problem, starch can be plasticized in combination with different synthetic polymers to satisfy a broad range of performance requirements of market needs.48–52 According to Bastioli,52 thermoplastic starch can be blended with synthetic polymers to generate three different categories of materials: •

Thermoplastic starch complexed with synthetic copolymers containing

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hydrophylic and hydrophobic units (i.e. copolymers of vinylalcohol, polyester-urethanes, ethylene-acrylic acid copolymers, etc.); Thermoplastic starch blended with incompatible synthetic polymers (cellulose derivatives, aliphatic polyesters, etc.) Partially complexed and/or compatibilized thermoplastic starch blended with incompatible or slightly compatible synthetic polymers.

An example of a biodegradable system based on starch that belongs to the first category includes the blend of starch with polyethylene-vinyl alcohol (EVOH). In this system, starch and the synthetic polymer form an ‘interpenetrated’ structure that results in total insolubility of starch and the biodegradation rate of starch is inversely proportional to the content of the amylose/vinyl alcohol copolymer complex.52–54 Other systems are based in incompatible synthetic polymers such as cellulose derivatives or aliphatic polyesters. Regarding aliphatic polyesters, the systems based poly-εcaprolactone and its copolymers have been subjected to intensive study by several authors.49,50,55–58 Starch was found to have a determinant effect on the biodegradation and mechanical performance of these type of blends.49,55,57 Other aliphatic polyesters blended with starch include polylactic acid (PLA),59,60 polybutylene succinate adipate (PBSA)48,50 and polyhydroxyalkanoates (PHAs).61,62 Starch can also be blended with cellulose derivatives like cellulose acetate, yielding an immiscible blend. Figure 3.1 presents a scanning electron

3.1 Scanning electron micrograph of a tensile failure surface of a blend of starch with cellulose acetate (SCA) featuring two distinct phases: a continuous (cellulose acetate) and a discontinuous (starch).

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micrograph of a tensile failure surface of a blend of starch with cellulose acetate (SCA), in which two distinct phases are clearly distinguishable: a continuous (cellulose acetate) and a discontinuous (starch).

3.3.1

Starch-based materials in the context of biomedical research

The improvement in physicochemical performance of starch-based blends as compared to TPS justified former studies as potential candidates in many different biomedical applications. In this context, Reis et al.63–100 have developed an extensive work concerning the investigation of several blends of corn starch (in amounts varying from 30 up to 50%wt) with several different synthetic polymers, namely: • • • •

Polyethylene-vinyl alcohol (EVOH), further referred as SEVA-C; Cellulose acetate, further referred as SCA; Poly-ε-caprolactone, further referred as SPCL; Polylactic acid, further referred as SPLA.

Several studies66,101–103 have shown that these blends degrade by both hydrolytic processes and enzymatic activity. The biomedical potential of these blends of starch is supported by the biocompatible character, which has been demonstrated in several in vitro81,89,99,104–107 and in vivo studies.105,107–109 The properties of these starch-based blends can be adequately tailored through the adequate selection of the synthetic component and the processing route. Depending on their synthetic component (polyethylene-vinyl alcohol, cellulose acetate, poly-ε-caprolactone or polylactic acid), these blends can present different mechanical behaviours ranging from an almost rubbery like material (SPCL) to a stiff one (SEVA-C, SCA or SPLA). In terms of processing, these blends can be processed as any ordinary thermoplastic by conventional melt based processing techniques. These blends of starch also exhibit a wide processing window as compared to standard TPS. Nevertheless, the starch fraction is still prone to thermo-mechanical degradation. So far, several processing methodologies have been employed with these materials, namely: extrusion compounding, melt spinning, compression moulding, and injection moulding. In the following sections, a brief overview of the processing of some starch-based blends is given which clearly illustrates the versatility of these materials.

3.3.2

The processing, structure and properties of starch-based blends

Mechanical performance of starch-based blends depends mostly on the blend composition. As an example, Figure 3.2 presents the evolution of fracture

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SCA

3.2 Evolution of fracture propagation upon impact testing at 0 and 2 ms for SCA: average impact mechanical properties: peak force (N) = 163.6 (17.3), peak energy (J) = 0.09 (0.01) and failure energy (J) = 0.23 (0.06). 0, 2, 4, 6 and 12 ms for SPCL: average impact mechanical properties: peak force (N) = 258.1 (10.9), peak energy (J) = 0.35 (0.04) and failure energy (J) = 0.95 (0.20).

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propagation upon impact testing for injection moulded samples of blends of starch with cellulose acetate (SCA) and poly-ε-caprolactone (SPCL). SCA is a rather stiff and fragile blend as compared with the ductile and tough character of the SPCL. Reis et al.63,64 originally proposed the use of starch-based blends as a bone-analogue material for the temporary fixation of bone fractures. In these studies, the processing of the blend of starch with polyethylene-vinyl alcohol (SEVA-C) by injection moulding for the production of compact specimens was described together with the respective mechanical performance characterization. The use of an alternative moulding technique, shear controlled orientation in injection moulding polymer (SCORIM), resulted in a significant improvement of stiffness and strength of SEVA-C as compared to conventionally moulded samples. In SCORIM, the polymer solidification takes place under a controlled macroscopic shear field that induces orientation of the molecular structure.110,111 In another work, Sousa et al.71 reported the enhancement in both stiffness and strength of SEVA-C as compared to conventional injection moulding. For conventionally injection moulded SEVAC, the typical values of tangent modulus and ultimate tensile strength are 2.2 GPa and 41 MPa respectively. SCORIM application results in a 31% increase in stiffness and a 19% increase in strength. The stiffness of SCORIM mouldings was found to depend on parameters such as holding and piston pressures during shear application, which define cavity pressure inside the mould. From these results, it is evident that SEVA-C can develop improved mechanical performance which appears to be correlated with the development of anisotropy inside the moulding upon SCORIM application. The development of preferred orientation in this semi-crystalline polymer is also possible at lower magnitude when promoting the melt over-flow inside the mould upon filling of the cavity. Over-flow conditions are attained when the mould cavity has an exit aperture that allows the flow of the material to continue after the filling of the cavity. Figure 3.3 presents a schematic diagram that compares the moulding geometries for conventional moulding, over-flow, and SCORIM. When SEVA-C is processed using these three different injection moulding approaches, conventionally moulded SEVA-C presents the lowest values of stiffness and strength, with 1.9 GPa and 40.4 MPa for respectively modulus and tensile strength. Over-flow mouldings exhibit 2.3 GPa of modulus and 45.5 MPa of tensile strength. The highest value of stiffness is presented for SCORIM with 2.5 GPa, even though the strength was found equivalent to over-flow conditions. When analysing the wide-angle X-ray diffraction patterns for these three mouldings (Figure 3.4), it is possible to characterize. SCORIM processed SEVA-C scattering by peaks at 2θ of 11.0, 12.8, 20.3 and 22.1°. The peak at 2θ of 11.0 and 22.1° are not observable for conventional moulding, while the last is observable for over-flow mouldings. Simmons and Thomas112 studied the structure of different starch/EVOH blends. Starches

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Relative intensity

3.3 Mouldings produced for SEVA-C: (a) conventional injection moulding, (b) over-flow; and (c) SCORIM.

SCORIM

Conventional moulding

Over-flow

5

10

15

20

25

30

35

40

2θ

3.4 Wide angle X-ray diffraction spectra for SEVA-C mouldings: conventional injection moulding, over-flow and SCORIM.

with amylose/amylopectin ratios of 0, 3/7 and 7/3 respectively were studied. For all cases, EVOH scattering peaks in starch/EVOH blends were reported at 2θ of 10.8, 20.4 and 21.7°. In native corn starches (amylose/amylopectin ratio of 3/7), the crystallinity of the starch fraction is related with the peak at 2θ of 12.8°. The remaining crystalline scattering of starch overlaps with the more intense EVOH scattering. The intensity of the peak at 2θ of 12.8° is

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similar for the three moulding sets in Figure 3.4. The distinction of crystalline peaks for both starch and EVOH in WAXD spectra suggests some degree of imiscibility between blend components. Simmons and Thomas113 also investigated the morphology and miscibility of corn starch/EVOH blends by transmission electron microscopy (TEM). Starch tends to form discrete domains along the EVOH matrix, even for high EVOH (70% by weight) contents. A limited degree of miscibility between starch and EVOH was also found to occur. The increase of intensity for EVOH peaks observed from conventional moulding<
3.3.3

The processing of starch-based composites

In the biomedical field, Reis et al.65,69 additionally proposed the development of anisotropic biodegradable composites through the combination of bioactive reinforcement of starch-based blends and the use of non-conventional processing techniques. In this approach, SEVA-C was combined with bioactive fillers and processed by means of SCORIM. Sousa et al.75,79 also explored this line of research by inducing an anisotropic character in injection moulded parts, through control of the structure development and adequate reinforcement strategy with hydroxyapatite (HA) particles. HA particles were employed in order to assure higher stiffness values as compared to standard stiffness values. For conventionally injection moulded SEVA-C/HA composites, it was possible to produce mouldings (with a rectangular cross-section of 8 mm2) with a modulus up to 6.5 GPa for a HA amount of 50% by weight (wt.). The use of SCORIM further extended the stiffness range of SEVA-C/ HA composites. For the same wt. amount of HA, it was possible to produce composite moulded parts with thicker cross-sections (circular cross-section with 20 mm2) and superior stiffness values – modulus of 7 GPa, which is in the bounds of the stiffness values reported for human cortical bone.114–116

3.3.4

The processing of tissue engineering scaffolds based on starch

While the previous research has focused mostly on the processing and characterization of starch-based blends on a compact form, our Group has

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also developed extensive research concerning the processing of porous starchbased blends to be used as scaffolds for tissue engineering research. Tissue engineering scaffolds have been produced by several melt based processing methodologies such as compression moulding combined with particulate leaching80 and extrusion or injection moulding combined with blowing agents.77,80,117 Using a processing method combining standard compression moulding and the use of leachable particles aimed at creating a porous network upon subsequent leaching, Gomes et al.80 developed starch-based scaffolds exhibiting pore sizes between 10 and 500 µm and porosity levels up to 50%. Conversely, foaming during melt extrusion or injection moulding can be obtained through the use of physical or chemical blowing agents that are responsible for the inducement of porosity. Tissue engineering scaffolds based on SEVA-C and SCA have been produced by melt extrusion employing endothermic chemical blowing agents, based on mixtures of citric acid and sodium bicarbonate.80 Using this method, scaffolds featuring porosity levels of 40–50% and pore sizes in the range of 100–500 µm were obtained. These scaffolds are biocompatible81 and exhibit adequate porosity and pore geometry for supporting cell growth and apparent bone formation.118 A similar strategy has been applied to the production of TE scaffolds based on injection moulding.77,117 However, injection moulding causes the formation of a compact outer layer on processed samples, which has to be subsequently removed. Using a completly different approach, our Group has also developed starchbased scaffolds119,120 based on melt spun SEVA-C, SPCL and SPLA fibres. In this method, fibres are first produced by melt spinning. For the case of SPCL, fibres typically present an average diameter between 200 and 300 µm (see Figure 3.5). Lower diameter fibres can be easily obtained by uniaxial deformation. However, structure and properties of the SPCL fibres change significantly upon cold deformation. For SPCL fibres, it is possible to obtain a concomitant increase in fibre anisotropy, which ultimately results in a stiffening effect. The increase in anisotropy upon uniaxial deformation is evident from the X-ray diffraction patterns presented in Figure 3.6. The reduced diameter and increased stiffness of cold drawn fibres could be used, in principle, to further enhance the porosity of scaffolds without adversely affecting their mechanical performance. Fibre mesh scaffolds can be finally produced by applying a heat treatment to bond random fibres bundles. By controlling the compression of the bundles, it is possible to obtain scaffolds with porosity values in the range between 50 and 80%. Depending on the morphology, these scaffolds exhibit compressive modulus between 0.3 and 21 MPa. In this case, mechanical performance does not depend solely on porosity. Other parameters such as average fibre length, contact area between fibres and average number of contact points between fibres are also assumed to play a role. Figure 3.7 presents micro-computer tomography for a fibre mesh scaffold based on SPCL.

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25

Frequency

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15

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5

0 50.00 100.00 150.00 200.00 250.00 300.00 350.00 Diameter (µm)

3.5 Scanning electron micrograph of a bundle of melt spun SPCL fibres (Diameter: 214 ± 50 µm).

3.3.5

The processing of starch-based blends by rapid prototyping

The difficulties and limitations in controlling scaffold morphology have led to the development of alternative processing methodologies based on rapid prototyping (RP). The main advantage is the possible integration of RP fabrication, computer assisted design (CAD) and medical imaging acquisition/ processing techniques for the production of anatomically adapted scaffolds featuring customized internal architectures. In the so-called 3D bioplotting, the material in the powder or granular form is heated inside a barrel and the molten material is displaced by a plunger or piston, while in precise extrusion the plasticization of the melt is made by a rotating screw. Our Group has been developing tissue engineering scaffolds based on SPCL using this approach. Figure 3.8 presents, as an example, a SPCL scaffold featuring a orientation pattern between consecutive layers of 0°/90°, a fibre thickness of 0.5 mm, a spacing between fibres of 1.0 mm (in the same layer). By controlling the scaffold architecture, it is possible to produce scaffolds with porosity values between 60 and 80%. As a consequence, these scaffolds exhibit compressive modulus between 0.3 and 23 MPa.

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(a)

(b)

3.6 Wide angle X-ray diffraction pattern of a SPCL fibre upon cold drawing: 30% (a) and 750% (b) uniaxial deformation.

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3.7 Micro-computer tomography obtained for fibre mesh SPCL scaffold (Porosity of around 65%).

3.8 SPCL scaffold produced by 3D bioplotting as observed in the isometric perspective featuring an orientation pattern between consecutive layers of 0°/90°.

3.4

Conclusions

Thermoplastic starch (TPS) is a thermosensitive material that degrades during processing. The application of TPS is rather limited by the inherent susceptibility to thermo-mechanical degradation. This limitation can be circumvented by combining starch with other synthetic polymers. In this chapter the processing and properties of blends of starch with polyethylenevinyl alcohol (EVOH) (SEVA-C), cellulose acetate(SCA), poly-ε-caprolactone

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(SPCL) and polylactic acid (SPLA) was described. These blends can be easily processed by different conventional and non-conventional routes into a large variety of forms and porous architectures. Mechanical performance is mostly dependent on blend composition. Nevertheless, processing is rather important as it can affect crystallinity and consequently mechanical performance. For porous morphologies, mechanical performance of scaffolds can vary significantly depending on the morphology of the scaffolds.

3.5

References

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76 Leonor I B, et al., Development of highly bioactive and mechanically strong starch thermoplastic/bioglass composite biomaterials, Key Engineering Materials, 2001, 192–195, 705–708. 77 Gomes M E, et al., A new approach based on injection moulding to produce biodegradable starch-based polymeric scaffolds: Morphology, mechanical and degradation behaviour, Biomaterials, 2001, 22(9), 883–889. 78 Vaz C M, Reis R L and Cunha A M, Use of coupling agents to enhance the interfacial interactions in starch-EVOH/hydroxylapatite composites, Biomaterials, 2002, 23(2), 629–635. 79 Sousa R A, et al., Mechanical performance of starch based bioactive composite biomaterials molded with preferred orientation, Polymer Engineering and Science, 2002, 42(5), 1032–1045. 80 Gomes M E, et al., Alternative tissue engineering scaffolds based on starch: Processing methodologies, morphology, degradation and mechanical properties, Materials Science and Engineering C, 2002, 20(1–2), 19–26. 81 Salgado A J, et al., Preliminary study on the adhesion and proliferation of human osteoblasts on starch-based scaffolds, Materials Science and Engineering C, 2002, 20(1–2), 27–33. 82 Leonor I B, et al., Novel starch thermoplastic/Bioglass® composites: Mechanical properties, degradation behavior and in-vitro bioactivity, Journal of Materials Science: Materials in Medicine, 2002, 13(10), 939–945. 83 Oliveira A L, Alves C M and Reis R L, Cell adhesion and proliferation on biomimetic calcium-phosphate coatings produced by a sodium silicate gel methodology, Journal of Materials Science: Materials in Medicine, 2002, 13(12), 1181–1188. 84 Alves C M, et al., Biocompatibility study of biodegradable starch-hydroxylapatite particulates for bone/dentistry fillers, Key Engineering Materials, 2003, 240–242, 725–728. 85 Mano J F, Koniarova D and Reis R L, Thermal properties of thermoplastic starch/ synthetic polymer blends with potential biomedical applicability, Journal of Materials Science: Materials in Medicine, 2003, 14(2) 127–135. 86 Sousa R A, et al., Processing and properties of bone-analogue biodegradable and bioinert polymeric composites, Composites Science and Technology, 2003, 63(3– 4), 389–402. 87 Altpeter H, et al., Shear controlled orientation in injection moulding of starch based blends intended for medical applications, Plastics, Rubber and Composites, 2003, 32(4), 173–181. 88 Boesel L F, Mano J F and Reis R L, Optimization of the formulation and mechanical properties of starch based partially degradable bone cements, Journal of Materials Science: Materials in Medicine, 2004, 15(1), 73–83. 89 Salgado A J, Coutinho O P and Reis R L, Novel Starch-Based Scaffolds for Bone Tissue Engineering: Cytotoxicity, Cell Culture, and Protein Expression, Tissue Engineering, 2004, 10(3–4), 465–474. 90 Mano J F, et al., Bioinert, biodegradable and injectable polymeric matrix composites for hard tissue replacement: State of the art and recent developments, Composites Science and Technology, 2004, 64(6), 789–817. 91 Boesel L F, Fernandes M H V and Reis R L, The behavior of novel hydrophilic composite bone cements in simulated body fluids, Journal of Biomedical Materials Research – Part B Applied Biomaterials, 2004, 70(2), 368–377. 92 Correlo V C, et al., Optimization of the injection moulding processing conditions

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94 95

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Natural-based polymers for biomedical applications of novel starch/poly(lactic acid) biomaterials, in Transactions – 7th World Biomaterials Congress, 2004, Sydney. Wang Y, et al., Thermal and thermomechanical behaviour of polycaprolactone and starch/polycaprolactone blends for biomedical applications, Macromolecular Materials and Engineering, 2005, 290(8), 792–801. Silva G A, et al., Starch-based microparticles as a novel strategy for tissue engineering applications, Key Engineering Materials, 2006, 309–311 II, 907–910. Ghosh S, et al., Osteochondral tissue engineering constructs with a cartilage part made of poly(L-lactic acid)/starch blend and a bioactive poly(L-lactic acid) composite layer for subchondral bone, Key Engineering Materials, 2006, 309–311 II, 1109– 1112. Boesel L F and Reis R L, The effect of water uptake on the behaviour of hydrophilic cements in confined environments, Biomaterials, 2006, 27(33), 5627–5633. Silva G A, et al., The effect of starch and starch-bioactive glass composite microparticles on the adhesion and expression of the osteoblastic phenotype of a bone cell line, Biomaterials, 2007, 28(2), 326–334. Alves N M, et al., Microhardness of starch based biomaterials in simulated physiological conditions, Acta Biomaterialia, 2007, 3(1), 69–76. Oliveira J T, et al., A cartilage tissue engineering approach combining starchpolycaprolactone fibre mesh scaffolds with bovine articular chondrocytes, Journal of Materials Science: Materials in Medicine, 2007, 18(2), 295–302. Ghosh S, et al., The double porogen approach as a new technique for the fabrication of interconnected poly(L-lactic acid) and starch based biodegradable scaffolds, Journal of Materials Science: Materials in Medicine, 2007, 18(2), 185–193. Azevedo H S, Gama F M and Reis R L, In vitro assessment of the enzymatic degradation of several starch based biomaterials, Biomacromolecules, 2003, 4(6), 1703–1712. Alberta Arauüjo M, et al., In-vitro degradation behaviour of starch/EVOH biomaterials, Polymer Degradation and Stability, 2001, 73(2), 237–244. Vaz C M, Reis R L and Cunha A M, Degradation model of starch-EVOH+HA composites, Materials Research Innovations, 2001, 4(5–6), 375–380. Marques A P, et al., Effect of starch-based biomaterials on the in vitro proliferation and viability of osteoblast-like cells, Journal of Materials Science: Materials in Medicine, 2005, 16(9), 833–842. Marques A P, Reis R L and Hunt J A, In vitro and in vivo inflammatory reaction to starch-based biomaterials, Cellular and chemical mediators, in Transactions – 7th World Biomaterials Congress, 2004. Sydney. Marques A P, Reis R L and Hunt J A, The effect of starch-based biomaterials on leukocyte adhesion and activation in vitro, Journal of Materials Science: Materials in Medicine, 2005, 16(11), 1029–1043. Santos M I, et al., Response of micro- and macrovascular endothelial cells to starch-based fiber meshes for bone tissue engineering, Biomaterials, 2007, 28(2), 240–248. Marques A P, Reis R L and Hunt J A, An in vivo study of the host response to starch-based polymers and composites subcutaneously implanted in rats, Macromolecular Bioscience, 2005, 5(8), 775–785. Salgado A J, et al., In vivo evaluation of extruded starch based scaffolds aimed for bone tissue engineering applications. in Transactions – 7th World Biomaterials Congress, 2004, Sydney.

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4 Controlling the degradation of natural polymers for biomedical applications H. S. A Z E V E D O, T. C. S A N T O S and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

4.1

Introduction

The structural similarity of some natural polymers with the components of the extracellular matrix (ECM), makes them interesting candidates as biomaterials. The present book describes the most important applications of natural polymers in the biomedical field, confirming the increasing interest on these polymers as biomaterials. The natural polymers with relevance in the biomedical area can be divided into three major classes: polysaccharides (alginate, hyaluronan, dextran, starch, cellulose derivatives, chitin derivatives), proteins (collagen, gelatin, fibrin, elastin, silk fibroin, soy protein) and bacterial polyesters (polyhydroxyalkanoates) (Mano et al., 2007, Hayashi, 1994).

4.2

The importance of biodegradability of natural polymers in biomedical applications

The demands for biomaterials with controlled degradation kinetics includes a wide range of biomedical applications such as resorbable surgical sutures, matrices for the controlled release of drugs, scaffolds for tissue engineering, wound dressing membranes, and resorbable devices such as bone cements, pins, screws and plates. In general, biomaterials used in these applications should degrade over time, so that completely natural tissue can develop without the long-term remainder of foreign elements, and finally clear from the organism once they complete their functions in the body. Additionally, the performance of many biomaterials depends largely on their degradation behaviour since the degradation process may affect a range of events such as cell growth, tissue regeneration, drug release, host response and also the material function. For example, degradability and degradation rate of the supportive matrix were identified as having a strong influence on cell migration, proliferation, differentiation and morphology of the newly formed tissue. For instance, Kong and co-workers (Kong et al., 2004a) showed that when 106 © 2008, Woodhead Publishing Limited

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alginate gels, with different degradation rates and containing primary bovine chondrocytes, were implanted into severe combined immunodeficiency disease (SCID) mice, tissues that were grown using more rapidly degrading alginate gels were largely more rigid than tissues from more stable gels. A histological evaluation of the engineered tissues also indicated that the degradation rate affected the cellularity and deposition of collagen (type II collagen and glycosaminoglycan) within engineered tissues, suggesting the formation of more mature cartilage tissue in the rapidly degrading gels. It was postulated that this effect was due to the fact that the accelerated degradation provided space that was essential for new tissue formation. In addition, it might be possible that in this case the host cells migrated into the rapidly degrading gels more effectively and participated in the tissue formation coupled with transplanted chondrocytes. It is important to note, however, that polymers which degrade too rapidly may not be able to serve as space filling scaffold capable of supporting new tissue development. Too rapid degradation may cause the matrices to collapse before substantial amounts of ECM were deposited by the cells. Consequently, it would be beneficial to couple the degradation to the rate of ECM production, in order to support tissue differentiation and tissue integrity. The optimal degradation rate will depend on the intended application as well the specificities of particular cells (metabolic activity of the cell type) in the host tissue or seeded in the tissue engineering construct. In addition to their biodegradability under mild conditions in the human body, the degradation products of natural polymers should be free of immunogenicity or toxicity and their absorption rates should be complete with the healing process. For temporary applications, the rate of degradation and the release of degradation products should be physiologically harmless, without inducing severe host responses and should be rapidly cleared from the organism. Several polymers are well accepted by the body, being nontoxic, but additives and degradation products can be cytotoxic. Thus, it can be seen that polymer degradation is an important factor in the suitability of polymers for biomedical applications.

4.3

Degradation mechanisms of natural polymers and metabolic pathways for their disposal in the body

Biodegradation is defined as ‘the gradual breakdown of a material mediated by specific biological activity’ (Ali et al., 1994). All polymers are susceptible to degradation but the conditions under which it occurs and the kinetics of the reactions are extremely variable. Degradation rate of biomedical polymers is mainly regulated by various intrinsic (average molecular weight, molecular weight distribution, crystallinity,

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hydrophilic-hydrophobic balance of the polymer, existence of hydrolysable bonds, surface area) and extrinsic (nature of the biological/cellular environment, applied stress) physico-chemical factors (Ali et al., 1994). The degradation processes of biomedical polymers can be generally divided into manufacturingderived degradation and degradation that is mediated by specific biological activity (biodegradation). During manufacturing, these polymers can undergo degradation when exposed to heat, shear stress, strong acids or alkali, radiation. Polymer thermal degradation may occur by pyrolysis (the chemical decomposition of polymers by heating in the absence of oxygen or any other reagents) or combustion (thermal degradation that occurs in the presence of high temperature and oxygen). Degradation caused by shear stress and high temperatures used during extrusion or injection moulding is an example of processing induced degradation. The thermal degradation of hyaluronan has been studied by thermogravimetric analysis (TGA) and infrared (IR) spectroscopy (Villetti et al., 2002). The TGA showed two mass loss stages, in which the first indicated the water loss and the second, the polysaccharide degradation (276°C). By IR spectroscopy it was observed that at low temperatures only scission of the exocyclic groups had occurred and the scission of the strong links in the backbone occurred at high temperatures. Degradation by radiolysis (cleavage of one or several chemical bonds resulting from exposure to highenergy flux) and photo-oxidation can occur during sterilisation steps (UV, γ and β radiations). The bioabsorption (absorption in the body) of polymeric materials can be divided into two distinct stages: degradation and absorption. The former stage involves breaking of bonds in the main chain, a decrease in molecular weight, and the production of oligomers and monomers. Under physiological conditions, natural polymers are mainly degraded by hydrolysis followed by oxidation (Hayashi, 1994). Hydrolysis is the scission of susceptible molecular groups by reaction with water (Figure 4.1). There are two different mechanisms for hydrolysis: polymers that are decomposed by enzyme-specific reactions (enzymatically degradable polymers) and polymers that are decomposed by contact with water or serum (nonenzymatically degradable polymers). Extracellular fluid components such as lipids are able to influence hydrolytic degradation of certain polymers. In addition, oxidative degradation of polymers may also take place, taking into account that several enzymes have oxidative activity and certain inflammatory cells release peroxide and other oxidative agents. After short term implantation, a large number of neutrophils and macrophages are recruited to the implant area. The phagocytic cells release cationic proteins and proteases and activated oxygen species (superoxide and hydrogen peroxide), major oxidative products of activated phagocytes (Sutherland et al., 1993). Polymer

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Polymer (insoluble) Monomer (soluble) Hydrolysable bond Water molecule

4.1 Schematic representation of polymer degradation by hydrolysis.

oxidation involves free radical chain reactions. The reaction is initiated by an oxidative specie followed by the propagation of the free radical by reaction with chain bonds. After being implanted in the body, biomedical polymers experience mechanical stress (loading, abrasion, wear, etc.) and these applied mechanical forces are expected to influence their degradation behaviour. It was shown (Ruberti and Hallab, 2005) that tensile strain directly modulates the susceptibility of collagen molecules to enzymatic degradation, i.e. it protects collagen molecules that are under load from cleavage. Fibrils under lower tensile load were preferentially cleaved. The degradation of polymeric biomaterials can occur in three different ways: polymers whose main chain bonds are hydrolysed to produce oligomers and monomers; polymers which are converted to soluble polymers by hydrolysis of side groups; and crosslinked polymers that are solubilised by dissociation of their cross-links. In the two latter cases, the molecular weight of the resulting soluble polymers should be low enough (< 50 kDa) to allow for ultimal renal clearance (Alsberg et al., 2003). Ionically cross-linked alginate hydrogels normally undergo slow dissolution, mainly due to the sensitivity of the gels towards calcium chelating compounds (e.g. phosphate,

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citrate and lactacte) or gradual exchange with monovalent cations present in the environment. They dissolve at neutral pH upon losing the divalent crosslinking cations, resulting in uncontrolled and typically slow degradation kinetics in vivo. Hyaluronan (HA) is enzymatically degraded by hyaluronidase and is completely resorbable through multiple metabolic pathways. Although HA is easily broken down in vivo, cross-linking individual HA polymer chains together decreases their degradation rate. Many strategies exist for cross-linking HA (see Section 4.5). Metabolic degradation of HA is principally intracellular. Although the turnover of HA has been established in several tissues, much less is known of its mode of disposal in these sites. It may be presumed that the presence of hyaluronidase is a prerequisite for metabolic degradation, and its distribution should indicate where the degradation may occur. The enzymatic degradation of HA in mammalian tissues takes places in two phases, including breakdown of the polysaccharide to its monosaccharide constituents and subsequent utilisation of the monosaccharide products. Degradation to the monosaccharide components is effected by the concerted action of three enzymes, hyaluronidase, β-D-glucuronidase and β-N-acetyl-D-hexosaminidase. Hyaluronidase is an endoglycosidase which cleaves internal β-N-acetyl-D-glucosaminidic linkages to D-glucuronic acid; β-D-glucuronidase is an exoglycosidase acting upon the non-reducing terminus of the oligosaccharides generated by hyaluronidase; and β-N-acetyl-D-hexosaminidase attacks the non-reducing terminal of β-Nacetyl-D-glucosamine residues resulting from the action of β-D-glucuronidase. Glucuronic acid and N-acetylglucosamine generated by the degradation of HA in the lysosomes diffuse into the cytosol, where they are metabolised further via the pathways (Figure 4.2). Fibrin is biodegradable being degraded within few days by the plasminogen in the culture medium (Jockenhoevel et al., 2001). The degradation of fibrin gels can be controlled and adjusted to the tissue development by the use of aprotinin (a protease inhibitor) which is able to stop the fibrinolysis via inhibiting plasmin completely or slow down the degradation of the gel in vitro. As a protein, silk is susceptible to degradation by proteolytic enzymes. Several authors (Arai et al., 2004; Hakimi et al., 2007; Horan et al., 2005; Minoura et al., 1990) had studied the biodegradation of silk fibroin films and fibres in presence of proteolytic enzymes (collagenase type F, α-chymotrypsin type I-S, protease type XXI). The rate and extent of degradation was dependent on the structural and morphological features of the polymer (fibre or film), enzyme-substrate ratio, type of enzyme and degradation time. Films were degraded more readily by the different enzymes although to a different extent. The protease type XXI was more aggressive towards silk films. Silk had shown, however, slow degradation in vivo. This could be valuable for applications that require support and transfer of load from the scaffold to the developing tissue.

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Hyaluronan (Hyaluronidase)

UDP-N-acetylglucosamine

Oligosaccharides

UTP

(Pyrophosporylase)

N-acetylglucosamine-1-phosphate (Glucuronidase) Glucuronic acid

(Hexosaminidase) (Mutase) N-acetylglucosamine

ATP

N-acetylglucosamine-6-phosphate

(Kinase) Gulonic acid

(Deacetylase) Glucosamine-6-phosphate

Acetate NH3 Carbon dioxide water

Acetyl CoA Oxidation Carbon dioxide water

(Deaminase)

Fructose-6-phosphate

Re-utilization Lipid, etc

Lactate

Carbon dioxide water

4.2 Pathways of hyaluronan catabolism in mammalian cells (adapted from Fraser et al., 1998).

Elastine-like peptides, with great potential to be used as an injectable system for cartilage repair, have been shown to be enzymatically degraded by trypsin (Ong et al., 2006). Extracellular collagen degradation occurs by the action of enzymes including metalloproteinases (MMPs) and through phagocytosis by macrophages and fibroblasts.

4.4

Assessing the in vitro and in vivo biodegradability of natural polymers

4.4.1

In vitro models

Generally, it is believed that the polymers known to degrade within the body are susceptible to hydrolysis and the in vivo degradation can be reproduced by simple aqueous solutions in vitro. In vitro studies enable one to predict in vivo degradation behaviour and are normally performed by incubating the materials in phosphate buffered saline solution (PBS, pH 7.4) alone or supplemented with relevant enzymes (Zhang and Cass, 2001; Tomihata and Ikada, 1997), in presence or absence of oxygen free radicals (Ali et al., 1994), at 37°C under static or dynamic conditions (agitation, flow). Different incubating media have been also used, like cell growth media (α-MEM,

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DMEM supplemented with 10% foetal bovine serum), human serum, inflammatory cells, cell lysates) (Rodriguez-Gonzalez et al., 2004; Tuovinen et al., 2004). Very often, the degradation tests are associated and performed in parallel with biocompatibility tests, both in vitro and in vivo. For instance, Chellat and co-workers (Chellat et al., 2000) have studied the degradation of chitosan-xanthan particles in complete culture medium for 1 and 28 days and then their extract products were assessed for cytotoxicity. In vitro degradation is examined by following the changes in the materials’ properties (surface morphology and chemistry, mass, mechanical properties, molecular weight, porosity) over time by using adequate characterisation techniques. For more details about degradation monitoring techniques please consult the reference by Azevedo and Reis (2005). Degradation products may be analysed by gel electrophoresis (SDS-PAGE), colorimetric methods, nuclear magnetic resonance (NMR), high performance liquid chromatography (HPLC) among others. For example, the soluble peptides formed during proteolytic degradation of silk fibres and films were analysed by high performance size exclusion chromatography (HP-SEC). Although in vivo degradation of biopolymers is a very complex process, involving various synergistic pathways of chemical, biochemical, physical and mechanical origin, in vitro degradation studies, using enzymes, are expected to contribute in elucidating the mechanisms by which polymers interact with the biological environment and characterise their functional properties. In fact, in the inflammatory process there is the direct involvement of circulating inflammatory cells recruited to the site of injury/implantation and the presence of serum is unavoidable. In addition, the cells with phagocytic ability are normally able to remove debris from the tissue by engulfment and digestion (Ali et al., 1994). The selection of the enzyme or enzymatic system is an issue that has to have in consideration the aim and specific use of the biomaterial. This is because, in different tissues, the most relevant enzymes may be also different. Another important topic to consider is the concentration of the enzyme or of each component into an enzymatic system. The concentrations should be selected according to the physiologic concentrations in each different tissue. Table 4.1 provides reference values for the concentrations of certain enzymes in the human serum. However, the enzyme concentrations vary from tissue to tissue and expression of certain enzyme activities increases during wound healing (e.g. early stages of implantation) and in several disease states. The effect of enzyme activities on the degradation rate of natural polymers has been studied. For instance, Burdick and colleagues (Burdick et al., 2005) studied the enzymatic degradation of photopolymerized hyaluronic acid networks by hyaluronidase and observed that the degradation is faster with higher enzyme concentration (100 vs 10 U/mL). The degradation behaviour of chitosan scaffolds was investigated in vitro

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Table 4.1 Reference serum concentrations and distribution of hydrolytic enzymes with relevance in the biodegradation of natural polymers in the human body Concentration

Occurrence in the human body

Potential substrate

Lysozyme (3.2.1.17)

Hydrolysis of 1,4-β-linkages between N-acetylmuramic acid and N-acetyl-Dglucosamine residues in peptidoglycan and between N-acetyl-D-glucosamine residues in chitodextrins

1.6 µg/mL (Selsted, 1978) 0.7–2.9 U/µL (Hollak et al., 1994)

Serous salivary acinar cells, lactating mammary tissue, paneth cells, renal tubular cells, myeloid cells (including eosinophils) and histiocytic cells (Mason and Taylor, 1975)

Chitin derivatives (e.g. chitosan)

Hyaluronidase (EC 3.2.1.35)

Random hydrolysis of 1->4-l inkages between N-acetyl-βD-glucosamine and D-glucuronate residues in hyaluronan.

2.6 U/mL (Delpech et al., 1987)

Bone marrow, testis and kidney (Fraser et al., 1998)

Hyaluronan, chondroitin and dermatan sulfates.

α-Amylase (EC 3.2.1.1)

Hydrolysis of α-1,4glycosidic linkages of starch in a random manner

46–244 U/L (Junge et al., 1989)

Pancreas and in salivary glands (Price and Stevens, 1999)

Acts on starch, glycogen and related polysaccharides and oligosaccharides

Lipase (EC 3.1.1.3)

Hydrolysis of triglycerides to partial glycerides and fatty acids. The pancreatic enzyme acts only on an ester-water interface; the outer ester links are preferentially hydrolyzed.

30–190 U/L (Chawla and Amiji, 2002)

Pancreatic acinar cells, digestive tract, adipose tissue, lung, milk and leukocytes (Tietz and Shuey, 1993)

Polyesters (e.g. Polyhydroxyalkanoates)

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Catalytic activity

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Enzyme

114

Enzyme

Catalytic activity

Concentration

Occurrence in the human body

Phospholipase A2 (EC 3.1.1.4)

Hydrolysis of the acyl ester bond at the sn-2 position of phosphoglycerides

1.3–10.8 µg/L (Nevalainen et al., 1992)

Pancreatic tissue and juice; inflammatory exudates. Expressed in various tissues such as macrophages, platelets, neutrophils, fibroblasts and lung endothelium (Nevalainen et al., 1992).

Collagenase, Matrix metalloproteinase 1, MMP-1 (EC 3.4.24.7)

Cleavage of the triple helix of collagen at about threequarters of the length of the molecule from the N-terminus, at 775-Gly-|-Ile776 in the alpha-1(I) chain

1.6–24 ng/mL (Manicourt et al., 1994)

Connective tissues (Bord et al., 1997)

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Potential substrate

Collagen

Natural-based polymers for biomedical applications

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by incubating the materials in a lysozyme solution and in vivo by implanting them subcutaneously in the back of rats (Wan et al., 2005). It was observed that the pore diameter of the scaffolds influenced their degradation behaviour both in vitro and in vivo. Another study showed that chitosan is degraded in vitro in the first 2 hours, by β-glucosidase, tested by loss of viscosity of the chitosan solution (Zhang and Cass, 2001), in a maximum of 6 h in solution with the enzyme. In this case the β-glucosidase was obtained from almond emulsin and it contains chitinase in its composition, otherwise the β-glucosidase would not be able to degrade chitosan, since that enzyme cannot degrade polysaccharides such as chitosan (Zhang and Cass, 2001). In general, chitosan with high deacetylation degrees degrades poorly, even in solutions with different concentrations of lysozyme (Tomihata and Ikada, 1997; Park et al., 2002; Alsberg et al., 2003; Wan et al., 2005; Freier et al., 2005; Zhang and Cass, 2001; Zhuang et al., 2007), being, in some cases, comparable to the degradation in PBS (Dallan et al., 2007). Those studies were always performed keeping in mind the particular use of the final biomaterial. Nonetheless, as mentioned before, the results can differ if different degradation systems are used for testing the in vitro degradability of the biomaterial (Dallan et al., 2007; Freier et al., 2005; Zhang and Cass, 2001; Shalaby and Park, 1994; Horan et al., 2005). Soy is a very resistant protein to the in vitro degradation and crosslinking with glyoxal did not change its behaviour in an isotonic saline solution buffered at pH 7.4 (Vaz et al., 2003; Mano et al., 2007). In general, proteins with a globular structure, such as soy, tend to be more resistant to hydrolysis than those with randomly coiled or helical structures (Vaz et al., 2003). Retinal Pigment Epithelium (RPE) cells can be considered active phagocytes and consequently rich in lytic enzymes into the lysossomal compartments (Tuovinen et al., 2004). Starch acetate microparticles were tested for degradation in the presence of RPE cells’ homogenates which were proven to contain the enzymes esterase and α-amylase and enhanced the degradation of the starch acetate microparticules (Tuovinen et al., 2004). Since the RPE cells have been widely used in replacement of animal models for testing the performance of intracorneal drug delivery systems, it is quite difficult to compare with the in vivo behaviour. Tablets of high-amylose corn starch were tested for drug delivery systems to be used in the gastrointestinal tract (Nabais et al., 2007). The tablets were tested in an in vitro release experiment in which the pH gradient was induced (from pH 1.2 to pH 7.4) to simulate the pH evolution into the gastrointestinal tract (Nabais et al., 2007). It was observed that a blend of starch with ethylenevinyl alcohol copolymer SEVA-C degrades faster immersed in an isotonic saline solution, than the same copolymer reinforced with 45S5 Bioglass® (Leonor et al., 2002). The same starch-based biomaterials were shown to be enzymatically degraded by α-amylase resulting in innocuous degradation

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products, such as glucose and other maltooligosaccharides (Azevedo et al., 2003). The Reactive Oxygen Species (ROS) are very important molecules contributing for degradation of natural polymers in vivo. The influence of ROS on the degradation behaviour of biomaterials can be tested in vitro (Kogan et al., 2007). The attachment of macrophages (the first cells in contact with the foreign material), to large areas such as implanted polymer surfaces will trigger phagocytosis, with consequent toxic oxygen metabolite synthesis and lysosomal enzyme release. The increase in synthesis of superoxide anion (O •2– ) and hydroxyl radical (•HO) as well as release of lysosomal hydrolases by activated macrophages will augment any tissue damage due to an injury (Ali et al., 1994). The reactivity of hydroxyl radicals is so great that, if they are formed in living systems, they will react immediately with whatever biological molecule is in their vicinity, producing secondary radicals of variable reactivity (Ali et al., 1994). It was demonstrated that reactive oxygen species (O •2– and •HO) produced by stimulated PMNs isolated by human blood, attack HA (Moseley et al., 1997) and the mechanism for the degradative effects of ROS on hyaluronan is thought to occur via a random modification to the unit monosaccharides, followed by the hydrolytic cleavage of the resulting unstable hydrolytic constituents, with a range of hydrolytic products chemically characterized (Uchiyama et al., 2000). Although it has been shown in vitro that different formulations of HA polymers (benzyl esterified HA and high molecular weight HA) tested with PMA-stimulated PMNs possess dose-dependent antioxidant properties against O •2– (with benzyl esterified HA having the greatest potential) (Moseley et al., 2003), low molecular weight HA showed the highest dose-dependent anti-oxidant properties towards •HO produced by PMNs (Moseley et al., 2003). Generally it is well accepted that the crosslinking of a natural polymer with itself or with other components induces higher resistance to degradation (Choi et al., 1999; Zhao, 2006; Liu et al., 2005). It was also shown that phosphate is an additive favouring the degradation of alginate (Polyzois and Andreopoulos, 1985). The degradation of alginate can be induced by irradiation (Nagasawa et al., 2000) and by either acid (Bouhadir et al., 2000) or enzymatic (Nakada and Sweeny, 1967) treatment, although it exhibits limited biodegradation (Shapiro, 1998). The degradation of Ca2+ crosslinked alginate gels can occur by the removal of the Ca2+ (Hollak et al., 1994) and, therefore, regulated by using chelating agents, such as lactate, citrate and phosphate (Hollak et al., 1994). This suggests that alginate gels degrade in vitro in phosphate solutions (Hollak et al., 1994). In vitro studies showed that esterified HA-based polymers with a lower percentage of esterification are more susceptible to depolymerization by hyaluronidase (Zhong et al., 1994). The variety of in vitro systems used to mimic the in vivo conditions had generated different

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results for the same material. Therefore, it is very important to keep in mind the limitations of each model and the considerations when designing the degradation experiment.

4.4.2

In vivo models

Biocompatibility and in vivo degradation are closely related in the sense that the degradation of a biomaterial implanted in a host is influenced by the presence and recruitment of inflammatory cells and consequently by the production and release of pro- or anti-inflammatory mediators; on the other hand, the degradation products of the implanted biomaterials, due to their physicochemical properties, may dictate the ongoing of the inflammatory process. The physiological environment of the human body can be aggressive to polymers. The degradation in some extent, as well as the kinetics and mechanisms of the process, can be significantly affected by several biological active species, especially by enzymes, lipids, peroxides, free radicals and phagocytic cells (Ali et al., 1994). Furthermore, the reaction of the host and the in vivo degradation may differ according to the local of implantation. The in vivo tests intend to predict what may happen in a human organism after the implantation of a biomaterial. They are considered as the last challenge of a biomaterial prior to clinical use. Mainly, in vivo degradation may occur either by enzymatic action and the enzymes usually are located inside the cell (namely phagocytic cells), or by hydrolysis (enzymatic or other) in the extracellular environment. The implantation of medical devices acting as foreign bodies leads to the development of inflammation which goes from an acute response to a chronic situation which, in the worst case, may lead to the rejection of the implant. The triggered responses derive from the activation of different complex mechanisms depending, in some extent, on the surface physicochemical characteristics of the implant. One of the important mechanisms is the activation of the Complement System, which can occur either by the classical (specific binding of antibodies to the surface) or by the alternative pathway (releasing specific molecules – C3b which is able to bind directly to the implant surface and C5a – leading to opsonization). The opsonization also leads to the release of oxidative metabolites. Since C5a is rapidly hydrolysed, it is important for leukocyte recruitment only in the acute phase of inflammation (Luttikhuizen et al., 2006). Degradation of implanted materials is closely related to the type of cells recruited and that persist in the local of implantation. Polymorphonuclear neutrophils (PMNs) and macrophages play a very important role in the progression of the response and degradation of the implanted materials, since their adhesion to the surface of the materials will dictate, to some extent, the formation of foreign body giant cells (FBGCs) and consequently the degradability. The development of the foreign body reaction (FBR) against implanted materials depends also

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on the physicochemical characteristics of the surface (Luttikhuizen et al., 2006). Usually, when a biomaterial is not biodegradable, after a period of time of in vivo implantation it will be surrounded by a fibrous capsule. In some cases, when this capsule around the implanted biomaterial is observed, it may mean that, from that point on, the implant will be not accessible to the cellular and enzymatic degradation action of cells and enzymes located outside the capsule. This capsule usually is formed when the body finds the implanted material harmful and tries to protect itself from eventual degradation products of the material. Depending on the purpose of the produced devices, the degradation test methodology should be adequate, even to the different tissues intended to be in contact with the polymer. In vivo degradation has been investigated by subcutaneous implantation of the polymeric materials into the backs of rodents and the implantation time can increase to 21 weeks. In vivo degradation studies of alginate films (Livnat et al., 2005) implanted in the subcutis of rats suggested that the main mechanism of degradation occurs via inflammation-mediated erosion of the material. Similarly to what happens in vitro, the crosslinking of natural polymers may increase the resistance to in vivo degradation (Jameela and Jayakrishnan, 1995; Chawla and Amiji, 2002; Liu et al., 2005). In the particular case of chitosan, higher deacetylation degrees induce lower degradation rates (Tomihata and Ikada, 1997). These results are in concordance of some in vitro results (Alsberg et al., 2003; Park et al., 2002), where the chitosan resisted the action of lyzozyme. A strategy to modulate the degradation rate of alginate is gamma radiation (Alsberg et al., 2003), which decreases the molecular weight of alginate in a dose-dependent manner. This was shown by measuring the degradation of alginate hydrogel discs after subcutaneous implantation in rats (Alsberg et al., 2003). Esterified hyaluronan derivatives show different in vivo degradation rates, either in subcutaneous implantation (Jansen et al., 2004) or in intramuscular implantation (Campoccia et al., 1996) in rats, maybe due to the hydration properties. Once again, at the time to choose the animal model to test the in vivo degradability of a biomaterial, it is very important to have in mind the final use of the device and the site where it will perform its function. When testing in vivo degradability, the local of implantation should mimic, as best as possible, the final local of implantation in the organism, since different in vivo environments may induce different degradation rates for the same natural polymer and with the same shape. An example is demonstrated by Azab and co-workers (Azab et al., 2006). Different degrading crosslinked chitosan hydrogels were implanted subcutaneously and intraperitoneally in rats. The degradation of the slow degrading hydrogels did not show to be different depending on the animal model and the hydrogel almost remained after 28 days of implantation. However, in the case of the fast degrading crosslinked

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chitosan hydrogel, the hydrogel was shown to degrade faster when implanted intraperitoneally when compared with the subcutaneous implant after 14 days (Azab et al., 2006). These results suggest that if the material has some degradation potential, the animal model of intraperitoneal implantation promotes a faster degradation compared with the model of subcutaneous implantation, probably because the intraperitoneal cavity is an open cavity endorsing the recruitment of higher amounts of inflammatory cells that release specific enzymes deeply involved in the degradation of implanted materials. Galassi and colleagues (Galassi et al., 2000) showed that the esterified hyaluronan derivatives used for skin regeneration and seeded with fibroblasts, were able to dissolute in different human wounds (two different clinical cases) after 4–5 weeks of implantation (Galassi et al., 2000). Since inflammation plays a very important role in the progression of in vivo degradation, a particular model to understand the influence of the persistent inflammatory response on in vivo degradation of a particular naturalbased polymer (hyaluronic acid – HA) was created. A surgical injury was inflicted to the animals 14 and 28 days after implantation of the HA hydrogels. The aim was to provoke an inflammatory process in order to observe the reaction of the implanted hydrogels. Approximately 20% of crosslinked HA gel was degraded within 10 days after the implantation, but the residual gel was relatively stable over the period of 100 days. Such limited degradation during the first few days was considered to be due to the inflammation caused by surgical incision for implantation (Kamimura et al., 2001). In order to generate a preliminary inflammation, surgical injury was made by incising the dorsum skin for additional 4 cm length and then suturing. It was possible to conclude that the crosslinked HA gel was degradable in response to inflammation although the gel was not affected under normal health conditions (Kamimura et al., 2001). During inflammation excessive amounts of superoxide anion radical (O •2– ) and nitric oxide (•NO) are formed, and together may contribute to the formation of peroxynitrite, which degrades some natural-based polymers (Kogan et al., 2007). Since the enzymes superoxide dismutase and catalase have an antioxidative action (Kogan et al., 2007), they may be used as controllers of degradation for natural polymers in a combined application. Hypobromous acid (HOBr) is produced during inflammation and it was recently shown that this can react with glycosaminoglycans, resulting in the degradation of some natural polymers (Rees et al., 2007).

4.4.3

Comparison between in vitro and in vivo models

Several in vitro systems have been designed to test the degradation of biomaterials. To better mimic the in vivo conditions in an in vitro system, researchers have been using a cocktail of enzymes at physiological

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concentrations. It is important to consider, however, that in pathological conditions (e.g. after implantation of a foreign body with the consequent inflammatory response) there is an increase in the concentration of certain enzymes. In this sense, it might be useful to perform in vitro inflammatory/ immune response assays to assess the expression of lytic enzymes and oxygen reactive species prior to degradation tests. Having taken into account this scenario and the complexity of biological systems (distinct cells, molecules and fluids, in different time schedules and concentrations, and in different stages of metabolism) it might be very difficult to fully characterize the in vivo conditions in vitro. Therefore, the use of animal experimentation to test the performance and biodegradability in vivo constitutes a necessary step prior to the pre-clinical stage. In fact, different results have been obtained during in vitro and in vivo studies for the same natural polymer. For example, the degradation rate in an in vitro system (with lysozyme) showed that chitosan-based scaffolds degrade faster in the first few weeks (Wan et al., 2005; Freier et al., 2005; Zhang and Cass, 2001) but, contrarily, in an in vivo system, the degradation is faster in the end periods of implantations (~10 weeks) (Wan et al., 2005). However, other studies showed the same degradation behaviour for the same natural polymers either in vitro or in vivo (Tomihata and Ikada, 1997; Alsberg et al., 2003; Park et al., 2002). Nonetheless, the results obtained from in vitro and in vivo tests may be combined to provide a better understanding about the degradation mechanism of biomedical polymers.

4.5

Controlling the degradation rate of natural polymers

Controlling the degradation rate of polymers has been one of the critical issues in general biomaterials research, and has been widely investigated. Ideally, matrix degradation would occur in temporal and spatial synchrony with the formation of new tissue. Adapting the degradation kinetics to the rate of tissue formation is, however, a challenging task. Several approaches, based on physical, chemical and enzymatic processes, have been investigated to control the degradation rate of natural polymers. An interesting approach has been the development of polymeric systems with a self-regulated degradation mechanism. In these systems, the degradation process is initiated and/or controlled under certain environment conditions or in response to tissue responses. It is known that when a foreign material is implanted, an acute inflammatory response occurs with the consequent pH decrease and secretion of hydrolytic enzymes at the implant site by inflammatory cells. Within this context, it may be useful to take advantage of the inflammation process by making the degradation of the material sensitive to certain enzyme activities (Figure 4.3). For instance, Martins and co-authors (Martins et al.,

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Inflammatory cells

Macrophage Lymphocyte Hydrolytic enzymes

Eosinophil Oxygen rective species Mast cell

4.3 System with degradation behaviour sensitive to the feedback provided by the cells involved in the inflammatory response.

2004) proposed a new concept to develop chitosan scaffolds with in situ pore forming capability based on the fact that chitosan is degraded in vivo by lysozyme, an enzyme which is present in the lysossomes of phagocytic cells. Materials to be used in some applications, such as hard-tissue replacement, must combine adequate mechanical properties with controlled biodegradability. The material should degrade while maintaining a specified minimum mechanical strength to support the formation of new tissue. It may be difficult to achieve the desired combination of degradation and physical properties in a single material. Conventional approaches to accelerate the degradation rate of biomaterials normally deteriorate their mechanical properties in parallel. Therefore, methods that allow the degradation rate of polymers to be modified without significant change in their mechanical properties could be broadly useful. In this context, enzyme encapsulation technology can be used to incorporate hydrolytic enzymes into the polymeric matrices and then provide systems with controlled degradation at desired sites and at specific rates. Controlled degradation by enzymatic means can introduce several advantages having taken into account the high specificity of enzymes for their substrates and also because enzyme activity can be regulated by environmental conditions (e.g. pH, temperature, presence of certain substances like metal ions, etc.). In addition, the degradation kinetics can be adjusted by the amount of encapsulated enzyme into the matrix. For instance, Goldbart and co-authors (Goldbart et al., 2002) developed an enzymatically controlled responsive

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drug delivery system consisting of a starch-based tablet incorporating a nonactive α-amylase and a protein. The enzyme reactivation was made by the presence of calcium ions (which are known to be essential for the enzyme’s tertiary structure and catalytic activity) from the medium, which causes the tablet degradation and the concomitant release of the protein. This enzymatically controlled degradable system is illustrated in Figure 4.4. Biomaterial degradability is a critical design criterion for achieving optimal tissue regeneration with cell transplantation. It is widely assumed that coupling the degradation rate of polymers used in cell transplantation carriers to the growth rate of the developing tissue will improve its quantity or quality. The in vivo kinetics of alginate hydrogel disassembly depends to a greater degree on the local ionic environment in the tissues than on the structural characteristics of the polysaccharide network. Therefore, several efforts have been made to tailor their degradation behaviour. The Mooney group has carried out extensive work aimed at developing various approaches for controling the degradation of alginate hydrogels. Controlling the material degradation via simple physical

Polymeric matrix

Inactive enzyme

Degraded matrix (soluble product)

Trigger molecule

Active enzyme

4.4 Enzymatically controlled degradable system: encapsulation of a hydrolytic enzyme that will degrade the polymeric matrix upon activation with a trigger molecule (e.g. ion).

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dissociation of polymer molecules may provide advantageous over chemical degradation. A new approach to regulate the degradation kinetics of ionically cross-linked gels via controlling the dissociation rate of the polymer chains was proposed by Kong (Kong et al., 2004a). They also developed alginate hydrogels with a range of degradation rates by γ-irradiating high molecular weight alginate to yield polymers of various molecular weights and structures. Decreasing the size of polymer chains increased the degradation rate in vivo. In addition, it was observed that the more rapid degradation led to dramatic increases in the extent and quality of bone formation (Alsberg et al., 2003). In other studies, they have shown that partial oxidation (with sodium periodate) and bimodal molecular weight (i.e. a mixture of high MW polymer and polymer tailored to have lower MW) were successfully combined to regulate alginate gel degradation (Boontheekul et al., 2005; Bouhadir et al., 2001, Kong et al., 2004b). The mechanism of the degradation was found to be mainly due to hydrolytic chain scission. Polymer cross-linking has been widely used as a strategy to control the degradation rate of biomedical polymers. Cross-links are covalent bonds linking one polymer chain to another by a cross-linking agent (Figure 4.5). Cross-linking inhibits close packing of the polymer chains, preventing the formation of crystalline regions. The restricted molecular mobility of a cross-linked structure limits the extension of the polymer material under loading and increases its degradation resistance. For instance, degradation studies of cross-linked gelatin-alginate sponges showed that the higher degree of cross-linking gave more resistance to collagenase digestion (Choi et al., 1999). Modification of HA can be achieved by chemical derivatisation and chemical cross-linking. Cross-linking involves the formation of a ‘bridge’ between HA chains. The cross-linking agents can be integrated into the formed network with different ‘bridge’ lengths. Cross-linking agents contain multiple functional groups which are reactive with HA, including glutaraldehyde, formaldehyde, epoxides, and divinyl sulphone. Carbodiimide has also been used as a watercondensing agent for the modification of HA. By selection of different cross-

Crosslinking agent

Polymer n

4.5 Schematic representation of polymer cross-linking.

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linking agents and controlling the degree of crosslinking, the biostability of HA can be improved, according with the type of bond formed between HA chains (Zhao, 2006). Therefore, the biodegradability of cross-linked HA can be controlled by either adjusting the concentration of stable bonds such as ether bonds and/or the incorporation of biodegradable bonds such as ester bonds. Chemical modification of polymer chains with hydrophobic molecules (Figure 4.6) has been used to reduce the polymer’s susceptibility to enzymatic hydrolysis. The most commonly used approach for derivatisation of HA is by esterification and amidation with the biodegradability being controlled by the adjustment of esterification degree and the selection of esterifiying agents (Band, 1998; Prestwich et al., 1998). On the contrary, the introduction of hydrophilic groups is expected to enhance the degradation of modified polymers. This approach has been applied to cellulosic materials which have shown poor degradation in vivo (Martson et al., 1999). Decreasing crystallinity and increasing hydrophilicity are known approaches to improve their biodegradability. Cellulose can be converted to different derivatives (carboxymethylcellulose, cellulose nitrate, cellulose acetate, cellulose xanthate) and some of these derivatives have been employed with success as biomaterials, showing gradual degradation over time (Entcheva et al., 2004; Miyamoto et al., 1989; Devi et al., 1986; Singh et al., 1982).

4.6

Concluding remarks

Although the interest in natural polymers for biomedical applications has increased significantly in recent years, the degradation mechanisms of most Hydrophilic groups

Biomaterial

Hydrophobic groups

Biomaterial

4.6 Typical approaches to control the degradation rate of polymers by chemical modification.

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of them are not completely elucidated. In fact, there is lack of information regarding the mode of disposal (metabolic pathways, excretion) of their degradation products in the human body. These studies are considered very important since the biochemical processes governing metabolism of degradation products largely determine their elimination and toxicity. The absence of suitable biological assays or other analytical methods for identifying and quantifying polymer degradation products in vivo obviously limits evaluation of their catabolism. Nevertheless, as knowledge about the degradation behaviour of these natural materials increases, new approaches for controlling their degradation rates are being developed for designing better materials to match desired integration rates in vivo and support the development of more functional tissues.

4.7

Acknowledgements

H. S. Azevedo acknowledges financial support under the EU funded Marie Curie Outgoing Fellowship Project POLYSELF (Contract No MOIF-CT2006-021948). This work was partially supported by the Portuguese Foundation for Science and Technology through funds from the POCI 2010 and/or FEDER Programmes, the European Union funded STREP Project HIPPOCRATES (NMP3-CT-2003-505758) and the European NoE EXPERTISSUES (NMP3CT-2004-500283).

4.8

References

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5 Smart systems based on polysaccharides M. N. G U P T A and S. R A G H AVA, Indian Institute of Technology Delhi, India

5.1

What are smart materials?

5.1.1

Smart water soluble polymers and smart hydrogels

Responding to changes in their external environment is a hallmark of living systems. On a larger time scale, evolution is believed to be adapting to the (changing) environment. On a shorter time scale, the design of living systems ensure that they not only cope up but adequately deal with what is happening around them (Vincent, 2000; Jeong and Gutowska, 2002). This is true at all levels of organization: species → organism → organ → tissues → cells → molecules. Biological molecules like proteins are considered very smart machines. This is reflected in their regulatory features such as allosteric regulation (Urry, 1993; Alberts et al., 1994). In recent years, two kinds of smart materials (Table 5.1) have attracted great attention: smart water soluble polymers and smart hydrogels. The former can change their solubility in a medium in response to one or more of a stimulus/stimuli (Table 5.2, Figure 5.1) (Roy and Gupta, 2003). The nature of these stimuli can vary: changes in pH, temperature, presence of a chemical species are the most common stimuli. Some other stimuli that have been used in this context are: electric field, solvents, light, and pressure (Kim and Table 5.1 Kinds of smart materials Smart polymeric materials

Nature of stimulus

Reference

Chitosan Alginates Carrageenans Methylcellulose Gellan Xyloglucan

pH Ion, pH Ion Temperature Ion Temperature

BeMiller (1965) Smidsrød and Draget (1997) Van de Velde and Ruiter (2002) Haque and Morris (1993) Masteiková et al. (2003) Kumar et al. (2002)

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Natural-based polymers for biomedical applications Table 5.2 Types of stimuli Smart polymer

Insoluble

Soluble

Chitosan

> pH 6.5

< pH 5.5

2+

Alginate

Ca < pH 2.0

EDTA > pH 2.0

Carrageenan

K+

Water

+ Stimulus Water molecules – Stimulus

Swollen polymer

Collapsed polymer

5.1 Smart polymeric material.

Park, 2002). These polymers are also called reversibly soluble-insoluble polymers. The second kind, smart hydrogels, change their shape/volume in response to similar stimuli (as used in the case of smart polymers) (Peppas, 1995; Hoffman, 2002). These changes are accompanied by uptake/release of large amounts of solvent. Such smart polymeric materials can be fashioned out of naturally occurring sources or can be synthesized using normal chemistry which is used for synthesizing polymers (Roy et al., 2004).

5.1.2

Polysaccharides as smart materials

Among the naturally occurring polymers which can be used as such as smart polymers or can be turned into smart hydrogels, polysaccharides constitute the most common and important molecules (Cascone et al., 2001). This chapter discusses some of the important polysaccharides. In each case, the ways to obtain these carbohydrates and the structural basis for smartness (with the nature of stimulus/stimuli identified) are briefly discussed. This is followed by a discussion on their various applications in the context of their smart behaviour. It is also brought out that the smartness is a seamless feature that runs through the various formats in which such materials are used. Such formats include tablets, films, microspheres and nanoparticles.

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5.2

Chitin and chitosan

5.2.1

Natural occurrence and purification

131

Chitin is a naturally occurring polyaminosaccharide. It occurs in the shells of crustaceans, exoskeletons of insects and cell walls of fungi. It is synthesized (and degraded) in the biosphere at the rate of >10 gt/yr which makes it an important renewable biomass. Commercially, wastes from the seafood processing industry constitute the source for chitin. Chitinases occur fairly widely and account for the biodegradable nature of chitin (Cosio et al., 1982). Alkaline N-deacetylation of chitin produces chitosan, which consists of ≥ 80–85% free amino groups. Chitin degradation in nature is quite slow. An estimate in 1999 showed that shell fish processing discards constitute 50– 90% of the total solid waste landing in USA. At the global scale, the estimate of this type of waste was 5.118 × 106 Mt/y. Shrimp and crab shell waste constitute the most widely used source for isolation and purification of chitin (Shahidi et al., 1999). It is also isolated from fungal mycelia. The purification protocol of chitin from seafood waste follows the sequence of steps shown in Figure 5.2. Chitin subjected to 40–45% NaOH deacylates and produces chitosan. The deacylation degree can vary but in order to produce soluble chitosan (at low pH), about 80-85% deacetylation is necessary. It may be noted that some deacetylation happens during extraction of chitin itself so any chitin would have some limited degree of deacetylated amino groups. Chitosan can be purified by solubilizing in acids followed by filtration. Spray drying of the filtrate produces the chitosan powder. Kuera (2004) has described a method for obtaining crosslinked chitosan directly from fungal mycelium. Of all the commercially produced polysaccharides (e.g. cellulose, dextran, pectin, alginate, agar, agarose, starch, carrageenans and heparin), chitosan is the only basic polysaccharide. Both chitin and chitosan are nontoxic with LD50 of chitosan being 16g/kg body weight (similar to salt or sugar!). Chitosan can be sterilized by any of the sterilization methods without affecting even its physical properties (Singh and Ray, 2000). The reactivity of free –NH2 group, nontoxic nature, biodegradability and sterilizability has resulted in numerous applications of chitosan in a variety of areas. Of the two, only Waste

dil NaOH

Deproteinization

dil HCl

Chitin

5.2 Protocol for purification of chitin.

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Decolorization

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chitosan shows smartness as a pH-sensitive polymer and hence this chapter will focus more on applications of chitosan based materials.

5.2.2

Structure

Chitin is constructed from units of N-acetyl-D-glucose-2-amine. These are linked together in β(1→4) fashion (in a similar manner to the glucose units which form cellulose) (see Figure 5.3a). In effect chitin may be described as cellulose with one hydroxyl group on each monomer replaced by an acetylamine group. This allows for increased hydrogen bonding between adjacent polymers, giving the polymer increased strength. Chitin does not dissolve in water. Chitosan is obtained by means of alkaline N-deacetylation of chitin (see Figure 5.3b). This is done by removing acetyl groups from some of the Nacetyl glucosamine residues, leaving exposed amine groups capable of attaining positive charges in aqueous solutions at low pH; hence, chitosan can be dissolved at low pH. This active amine group provides many unique chemical and physical properties to the chitosan polymer.

5.2.3

Smart behavior of chitosan

Chitosan contains free amino groups with pKa ≈ 6.5. Hence at pH < 6.5, chitosan chains carry enough positive charge. This positive charge makes O

O

CCH3 NH O

H

CCH3 CH2OH

H OH H

O

O

H O

CH2OH

O

OH

H

H

NH

NH

H

H

OH

O

H

O

n

CH2OH

CCH3 O (a) NH2 O

H

CH2OH

H

O

OH H

H O

O CH2OH

O

OH

H

H

NH2 (b)

5.3 (a) Structure of chitin; (b) chitosan.

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NH2

H

H

OH

O

H CH2OH

O

n

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chitosan a cationic polyelectrolyte, which is soluble in water (and in dilute solutions of many organic acids such as formic acid, acetic acid, tartaric acid, citric acid). If the pH of an aqueous solution of chitosan is raised to above 7.5, the polymer precipitates as all the amino groups have ionized (–NH 3+ → –NH 2 ) and the polymer carries no charge. This process, being a simple ionization is reversible and makes chitosan a reversibly solubleinsoluble polymer or a pH responsive smart polymer (Terbojevich and Muzzarelli, 2000). Many methods of preparing chitosan hydrogels have been described (Draget et al., 1992). Such gels include chitosan-oxalate gels, chitosan-naphthalene sulphuric acid gels and gels prepared by crosslinking chitosan with Mo (+6). In most of these cases, unfortunately the chemistry of preparation is less than clear. The chitosan-Mo gels were found to swell nine times when placed in distilled water, the swelling capacity decreasing to 0 in 100 mM sodium chloride solution.

5.2.4

Applications

Chitin and chitosan constitute one of the most widely studied polymers from an application point of view. The application areas include waste water treatment, food industry, agriculture, paper and pulp industry, cosmetics, medicine, tissue engineering, bioseparation and biocatalysis (Dutta, 2005). It is not possible to cover all these applications here. The following overview of application focuses mostly on those applications which exploit the smart behavior of chitosan. Applications in enzymology One of the early applications of the smart behavior of chitosan is in the area of bioseparation of a lectin from wheat germ (Senstad and Mattiasson, 1989). Lectins are proteins of nonimmune origin that recognize free carbohydrate, or as part of glycoconjugates, in a specific fashion. This property makes these molecules as excellent tools in biology (Liener et al., 1986; Van Damme et al., 1997). The lectin from wheat germ is specific for N-acetylglucosamine. When chitosan solution was added to a crude homogenate of wheat germ, the polymer selectively complexed with the lectin. The ‘affinity complex’ could be precipitated by raising the pH and the lectin recovered after dissociation from the complex. This approach, called affinity precipitation, is a powerful tool in downstream processing of proteins/ enzymes (Gupta and Mattiasson, 1994). More details of this and other bioseparation techniques mentioned here can be found in Chapter 2. Subsequently, the similar approach was followed for purification of lectin from tomato and potato as these lectins also have similar specificity (Tyagi et al., 1996). The same principle was

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extended to developing an interesting version of a bioseparation technique called aqueous two-phase affinity extraction (Walter and Johansson, 1994; Hatti-Kaul, 2000). The technique exploits partition of a protein in a twophase polymer/polymer or polymer/salt system. PEG/salt is a frequently used two-phase system. One of the key constraints has been that it is difficult to separate partitioned protein from PEG. Incorporation of chitosan in the PEG phase, not only enhances the partition of protein (having binding affinity towards chitosan), affinity precipitation of the affinity complex from the PEG phase leaves the latter free for reuse (Teotia and Gupta, 2001a, b). In both approaches, the application of chitosan as a smart polymer can be extended beyond proteins which recognize chitosan (Mondal and Gupta, 2006). Apart from free –NH2 group, chitosan also has numerous hydroxyl groups. These two functionalities are valuable for linking any affinity ligand to chitosan. As the density of these affinity ligands on the polymer can be controlled and generally is not very high, such conjugation does not abolish the smart behavior of chitosan. It is possible that the pH of phase transition may change somewhat. The macro-(affinity ligand) so synthesized can be used either in affinity precipitation or aqueous two-phase extraction. As larger numbers of affinity ligands are available (Gupta, 2002), this creates a vast opportunity for chitosan in the area of bioseparation. Today, powerful technologies exist by which peptide libraries can be created for obtaining an affinity ligand for practically any enzyme/protein (Mondal and Gupta, 2006). This creates unlimited scope for chitosan (and similar materials) to be used in bioseparation. The smart behavior of chitosan can also be used to design smart biocatalysts. Enzymes as biocatalysts are superior to chemical catalysts as these proteins can act at normal temperature and pressure and show high specificity. It is now also known that apart from aqueous milieu, enzymes can also function in neat solvents, reverse micelles and gaseous phase (Gupta, 1992; Gupta, 2000). One factor, which has limited their application, has been cost. Immobilization is a well-established technique for converting enzymes into reusable catalysts and consists of adsorbing, entrapping, encapsulating or covalently linking enzymes to polymeric matrices (Cao, 2005; Guissan, 2006). Conventionally, these polymeric matrices are insoluble materials like agarose or polyacrylamide. The concept works well except that as most of the enzyme molecules are within the polymeric network, the ‘mass transfer limitation’ is especially severe for macromolecular substrates. Considering that most of the biomass is macromolecular in nature, immobilized enzymes have not shown good performance in the area of biomass conversion. When the biomass is insoluble like lignocellulosic material, this conventional heterogeneous biocatalyst design is not much use. Smart polymers like chitosan as watersoluble matrices for enzyme immobilization provide an interesting option in the biocatalyst design. An enzyme linked to chitosan can operate at pH

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below 6.3 as a soluble (homogeneous) biocatalyst. After the reaction is over, the biocatalyst can be recovered by raising the pH and reuse can be evaluated (Roy et al., 2004). It is interesting to note that while both chitosan and chitin based insoluble matrices have been extensively used for enzyme immobilization (Krajewska, 2004), there is only one application of using chitosan for designing a smart biocatalyst. Laccase from Coriolopsis gallica was linked with chitosan via carbodiimide coupling. The conjugate showed reversible soluble-insoluble behavior. The immobilized enzyme had enhanced stability at both pH 1 and pH 13. This successful design should encourage use of chitosan as a smart matrix for obtaining smart biocatalysts for hydrolysis of macromolecular substrates. Pharmaceutical, biomedical and miscellaneous applications of chitosan Chitosan forms gels at low pH range and is reported to have antacid and antiulcer activities in the stomach. Both physical gels and chemically crosslinked gels are degraded by lysozyme and this allows the design of enzyme degradable hydrogels for drug delivery purpose. Chitosan malate granules as carriers have been reported to work well for sustained release effects for drugs. As these granules do not dissolve at the acidic pH of the stomach this is a cost-effective way of prolonging residence time of drugs in the stomach since drugs are shielded from deactivating enzymes and the acidic pH (Henriksen et al., 1993). Numerous studies related to this application have been reported (Singh and Ray, 2000). In case of injured tissues, chitosan and its derivatives help blood coagulation and accelerate wound healing. Chitosan implants in the cornea are reported to encourage neovascularization (Singh and Ray, 2000). Chitooligosaccharides and chitosan lactate have been shown to be useful in replacing other chemical preservatives for processed food materials. Chitosan films have been used as food wraps (Shahidi et al., 1999). Extended shelf life has been reported by the use of chitosan films in the case of fruits, vegetables and fish. The cationic nature of chitosan at pH 4.5 results in its complexing (and precipitating) milk fat globule fragments. This constitutes an industrially viable process for removing fat from whey (Shahidi et al., 1999). Chitosan as a food additive is reported to possess antioxidative and hypocholesterolemic effects (Shahidi et al., 1999). The interaction of this positively charged polymer with negatively charged skin and hair forms the basis of its usefulness in skin care and health care products. For the former application, it is also used as a matrix for minerals, liposomes, fragrances and pigments. Its films or gels on its own and with cross-links with anions have moisture retaining capacity which is valuable in skin care applications.

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Later discussion on composite materials will discuss how the stimuliresponsiveness of chitosan as a part of composite materials leads to some further very interesting applications.

5.3

Alginates

5.3.1

Natural occurrence and purification

Alginic acid occurs as the main cell wall constituent of brown macro algae in the form of mixed salts of Na+, Mg2+, and Ca2+ ions. Apart from these sea weeds, alginates are produced by the microorganism Azotobacter vinelandii and some Pseudomonas strains. Commercially available alginates are mostly isolated from Laminaria hyperborea, Macrocystis pyrifera and Ascophyllum nodosum. Some other minor sources are Laminaria digitate, Laminaria japonica, Eclonia maxima, Lesonia negrescens and Sargassum sp. (Smidsrød and Skjåk-Bræk, 1990). The soluble alginic acid is extracted from algae with 0.1–0.2 N mineral acid. Mechanical treatment of the suspension is necessary to facilitate diffusion of alginic acid out of the algal mass. This step removes other salts and polymers. Sodium alginate is obtained by neutralization with sodium hydroxide. The alginate is precipitated by the addition of CaCl2 or ethanol (Smidsrød and Draget, 1997). Polyphenols are present in most of the alginate preparations. While these contaminants cannot be removed completely, some of the applications for alginate require that their level is brought down to less than a few percent. Bleaching with H2O2 and NaClO2, repeated precipitation with ethanol or acetone and treatment with activated carbon or polyvinylpyrrolidone help in removal of polyphenols. Their presence can be evaluated by fluorescence spectroscopy (Skjåk-Bræk et al., 1989). Samples with different chain length of the polymer can be prepared by ultrasonication (Martinsen et al., 1989). The total worldwide production of alginates has been estimated to be around 30 000 Mtons per year. The algae are mostly harvested from cold and temperate waters of North Europe, South American west coast, Southern Australia, Japan, and China (Smidsrød and Draget, 1997). Alginates show polydispersity with respect to average molecular weights which are generally in the range of 50–500 kDa.

5.3.2

Structure

Alginates are linear unbranched polymers containing β (1→4) linked Dmannuronic acid (M) and α (1→4) linked L-guluronic acid (G) residues (see Figure 5.4). These monomers occur in the alginate molecule as regions made up exclusively of one unit or the other, referred to as M blocks or G blocks, or as regions in which the monomers approximate an alternating sequence.

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Smart systems based on polysaccharides COOH C

O

O

H

H

C

C

OH C

OH

OH

C

C

H

H

H

COOH C H

H C

H

C H

C H

O H

C O

O

H

OH

COOH

137

C O

C OH

OH

C

C

H

H

H

5.4 Structure of alginate.

The NMR demonstrated that ring conformations were 4C1 for mannuronic acid and 1C4 for guluronic acid. The D-mannuronic acid exists in the 1C conformation and in the alginate polymer is connected in the β-configuration through the 1- and 4- positions. The L-guluronic acid has the 1C conformation and is α (1→ 4) linked in the polymer. Because of the particular shapes of the monomers and their modes of linkage in the polymer, the geometries of the G block regions, M block regions, and alternating regions are substantially different. Much of the early work (in the 1960s) on alginates and their applications should be credited to the Norwegian Institute of Seaweed Research. Schematically, a typical alginate would look like: M-M-M-M-M……..M-G-M-G-M-G……..G-G-G-G-G…….. M block MG block G block It was found that the ratio of total M/total G is different in different species. Ascophyllum nodosum alginate has M/G = 2.7 whereas alginate from Laminaria hyperborea shows the extreme of M/G = 0.6. Interestingly, young tissues are rich in M blocks and the percentage of G blocks increases as tissue grows older. The bacterial alginates show the presence of O-acetyl groups. Interestingly, A. vinelandii initially produces poly M and extracellular enzyme mannuronan C-5 epimerase converts some M into epimer C-5 guluronic acid; O-acetyl groups wherever present inhibit epimerization. It has been shown that algal alginate’s composition can also be modified by this epimerase (Smidsrød and Draget, 1997).

5.3.3

Smart behavior of alginates

The pKa values for –COOH groups in M and G are 3.38 and 3.65, respectively. This results in precipitation of the soluble polymer below pH 2. However, most of the applications of alginate arise from the fact that it forms insoluble gels/precipitates with divalent metal ions, especially Ca2+. The affinity order for alginates is (Smidsrød and Skjåk-Bræk, 1990):

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Pb2+ > Cu2+ > Cd2+ > Ba2+ > Sr2+ > Ca2+ > Co2+ = Ni2+ = Zn2+ > Mn2+ Some multivalent ions like Ti3+ and Al3+ stabilize Ca2+-alginate gels. Running a sodium alginate solution as drops into a CaCl2 solution gives rise to fairly spherical beads of Ca-alginate. If another species like a drug, protein or cell is added to the sodium alginate solution, the species is entrapped in Caalginate beads. This has been exploited in a large number of applications related to drug release systems and whole cell immobilization (Smidsrød and Skjåk-Bræk, 1990). As chelators like EDTA can remove Ca2+ easily, alginates can be considered as Ca2+-responsive polymers. Common buffers like phosphate or citrate also chelate Ca2+. As Na and Mg alginates are soluble, these ions are called antigelling ions. It is necessary to keep Na+:Ca2+ ratio below 25:1 for G-rich and below 3:1 for M-rich alginates. High G-alginates result in Ca-alginate beads which are more porous, have higher mechanical stability and greater tolerance to salts and chelating compounds. These beads also show minimum volume change on swelling-deswelling (drying and resuspension in aqueous solutions) as compared to beads made from low G-alginates (Smidsrød and Skjåk-Bræk, 1990; Smidsrød and Draget, 1997). A good discussion on the fine structure of alginate gels is given in an excellent overview by Smidsrød and Draget (1997).

5.3.4

Applications

Alginate is a nontoxic biocompatible polymer and food grade alginate preparations are easily available. Thus, it is not surprising that this polymer is also used widely for numerous applications. In food materials, most of the uses of alginate originates in enhancing the viscosity. As a natural cold soluble hydrocolloid, it is used as a stabilizer/thickener in low fat margarines/ low fat spreads, salad mayonnaise/dressing, beverages, bakery products, ice creams, pet foods and restructured food (e.g. pimento fillings for olives!). Alginate (and its blends) are also used in jams, marmalades, textile printing, paper coating and as lubricants and binding agents in welding rod coatings (http://www.fmcbiopolymer.com/PopularProducts/FMCAlginates/Introduction/ tabid/795/Default.aspx). In biotechnology, the major applications of alginate involve its use as a material for entrapment. For whole cell immobilization, the simplicity of entrapment protocol has made Ca2+-alginate a favorite choice (Smidsrød and Skjåk-Bræk, 1990). For entrapment of other molecules of smaller size, e.g. drugs and proteins, composite materials containing alginate have been used more often (see Section 5.6 on composite materials). Let us look at the applications of alginate which directly exploit the smart nature of alginate. Alginate shows inherent selectivity in binding to quite a

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few enzymes. This possibility has turned alginate into a very valuable polysaccharide in the area of bioprocessing in general and bioseparation in particular. The applications of alginate in bioseparation of proteins was reviewed recently (Jain et al., 2006). Hence, only a summary of the results will be provided. An affinity complex of the target enzyme with alginate can be precipitated by Ca2+ from the crude protein extracts. As already discussed (in Section 5.2 on chitin and chitosan), this strategy is known as affinity precipitation. The enzymes which have been purified by this simple elegant method include pectinase, lipase, α-amylase, β-amylase, pullulanase and phospholipase D (Jain et al., 2006). A recent interesting observation is that microwave pretreatment of alginate resulted in higher selectivity of the polymer towards pectinase and 20-fold purification (as compared to 10-fold purification observed by using untreated alginate). Similarly, just as described for chitosan, alginate can also be incorporated into PEG–salt two-phase systems and used for purification of enzymes by affinity partitioning. Again, the affinity complex of alginate-target enzyme can be separated by exploiting the Ca2+-responsive property of alginate. As the conjugation chemistry with alginate is already available (Draget et al., 1988), it is easy to link any affinity ligand to alginate and extend its use as a soluble affinity material for other large numbers of enzymes and proteins (for a discussion on the use of affinity based separations, see for example, Gupta, 2002). The concept of smart biocatalyst design has already been discussed in the context of chitosan. The first such design was in fact reported with alginic acid (Charles et al., 1974). The polymer was linked to lysozyme. Later on Dominguez et al. (1988) covalently coupled β-galactosidase with alginate but no details of its catalytic performance for lactose hydrolysis were unfortunately provided. One reason why alginate has not been used more for smart biocatalyst design is that charged matrices (like alginate) bind a lot of proteins and other molecules (substrates/products) nonspecifically by electrostatic interactions. Even then, considering that synthetic polymers like methacrylates have been used quite extensively in smart biocatalyst design (Roy et al., 2004), alginate in that respect may be an underexploited polymer. Pharmaceutical applications Fathy et al. (1998) have used tiaramide, a nonsteroidal antinflammatory drug (with a short half life), in alginate beads as a sustained release formulation. Pharmacokinetic parameters measured during in vivo experiments showed that high G alginate gave the best results. Earlier, Downs et al. (1992) described a slow release system for growth factors and concluded that entrapment in alginate beads constitutes an effective localized and slow release delivery system for biologically active molecules. Bodmeier et al. (1989) exploited

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the fact that Ca2+-alginate beads remained intact in 0.1 N HCl but dissolved in intestinal fluids to develop an oral formulation for delivery of micro- and nanoparticles as drugs. Alginate has been used fairly extensively in tissue engineering. Wang et al. (2003) showed that Ca-alginate is a good substrate for rat marrow cell proliferation. Yang et al. (2002) evaluated galactosylated alginate as a scaffold for hepatocyte attachment. It was shown that tissue engineered cardiac graft consisting of cardiomyocytes in alginate scaffold prevented damage after myocardial infarction in rats. The process for cardiac cell seeding and distribution in 3D alginate scaffolds has also been optimized (Dar et al., 2002). Alginate as a pseudochaperonin Alginate has been found to be a good additive for facilitating protein refolding (Mondal et al., 2006). Normally it is believed that polymers bind to hydrophobic patches in unfolded protein to prevent aggregation. The success of alginate as ‘pseudochaperonin’ indicates that polyelectrolytes may also serve the purpose by interacting with some charged residues in the unfolded protein.

5.4

Carrageenans

5.4.1

Natural occurrence and purification

Unlike land plants, marine algae produce large amounts of sulphated polysaccharides. The family of sulphated polysaccharides called carrageenans is one such class. More than 600 years ago, in the village of Carraghen situated on the south Irish Coast, flans were made by cooking the Irish moss (red seaweed species, Chondrus crispus) in milk. The use of Irish moss polysaccharides as a thickner, textile sizing and beer clarification has been mentioned (Velde and Ruiter, 2002). Commercial production began in the 1930s in USA when purified carrageenans were produced (Van de Velde et al., 2002). This family of polysaccharides is today produced from genus Chondrus, Eucheuma, Gigartina, and Iridaca of red seaweed Rhodophyceae. Purification A typical flowsheet for purification of carrageenan is shown in Figure 5.5 (http://www.fmcbiopolymer.com/PopularProducts/FMCCarrageenan/ Introduction/tabid/804/Default.aspx). The concentrated carrageenan solution is either converted into gel by running into KCl solution or precipitated by adding isopropyl alcohol.

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Harvest the seaweed, quickly dry it and bale it

Mechanical grinding and sieving to remove impurities (e.g. sand and salt)

Extensive washing

Hot extraction to solubilize carrageenan

Centrifugation to remove dense particles and filtration to remove small particles

Evaporation of water

5.5 Flowsheet for purification of carrageenan.

5.4.2

Structure

Carrageenan is a mixture of linear polymers of sulfated galactans which constitute cell wall material of marine red algae. Mostly, the chain consists of alternating units of 3-linked-β-D-galactopyranose (G-unit) and 4-linkedα-D-galactopyranose (D-unit) or 4-linked 3,6-anhydrogalactose (DA-unit). Other carbohydrate residues like xylose, glucose, their uronic acids and some other groups like pyruvic acid and methyl ethers (as substituents) are also present. The sulphate content is in the range of 22%–38% (w/w) (Van de Velde et al., 2002; Michel et al., 2006). Any algal extract contains a mixture of structural variants of carrageenan, the chemical structure depending upon the algal source and even the life stage of the algae and extraction procedures. The earliest investigations classified the carrageenans based on their solubility in KC1 solution as κ-carrageenan (insoluble) and λ-carrageenan (soluble). Later, through a vast amount of studies using various chemical and instrumental techniques such as alkali-treatment, methylation, partial acid hydrolysis, enzymic degradation and 13C-NMR and IR spectroscopy, this has been replaced, for the most part, by classification based on chemical structure. As a result, the carrageenans are divided into three families according to the position of sulfate groups in the 1,3- and 1,4-linked galactose residues. This classification is in terms of the nature of the repeating disaccharide made from D/DA and G units. The carrageenan preparations are called κ, ι, and λ corresponding to one, two and three sulphate groups per disaccharide (see Figure 5.6) (Michel et al., 2006). The presence of substituent groups, replacing hydroxyl groups, or other modifications of this disaccharide unit, such as anhydride ring formation, gives rise to the structural variants present in carrageenans. Therefore, carrageenans are a mixture of structurally related

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Natural-based polymers for biomedical applications (a) Kappa (κ) CH2OH

CH2 O

–O

3SO

O

H

O

H

O H

H H

H

H

O H

OH

OH

H

(b) Iota (ι) CH2OH

CH2 O

–O

3SO

O

H

O

H

O H

H

H

H

H

O OH

H

H

OSO3–

(c) Lambda (λ) CH2OSO3–

CH2OH O

O H

O

H

H

H H

H H

H

O H

OSO3–

H

OSO3–

5.6 Structure of carrageenans.

polysaccharides differing primarily in the proportions of galactose, ester sulfate (also in the position and content) and 3,6-anhydro galactose depending upon the species of carrageenophytes. κ-carrageenans are soluble in hot water, sodium ι-carrageenan is soluble in cold and hot waters and λ-carrageenan is partially soluble in cold water and completely soluble in hot water. The IUPAC names for κ, ι, and λ-carrageenan are carrageenose 4′ sulphate, carrageenose 2,4′ sulphate, and carrageenan 2,6,2′ trisulphate, respectively (Van de Velde and Ruiter, 2002). Commercially available food grade carrageenans have average molecular weight in the range of 400–600 kDa.

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5.4.3

143

Smart behavior of carrageenan

κ-Carrageenan forms gels with helical structures in the presence of K+ ions. Other monovalent ions which induce gel formation of κ-carrageenan solutions are Rb+, Cs+ and NH +4 (van de Velde et al., 2002). Ca2+ produces more compact and brittle gels. ι-Carrageenan forms dry and elastic gels with Ca2+ whereas λ-carrageenan forms free flowing, non-gelling, viscous (pseudoplastic) solutions in water. Gels prepared with κ- and ι-carrageenan are thermoreversible. These can be melted upon heating and reset upon cooling. Increase in concentration of the respective cations increases gelling temperature. It has been shown that 0.3% κ-carrageenan at about 38°C is about >75% precipitated with 0.2% KCl. The precipitate could be dissolved in distilled water. Hence, in a way, κ-carrageenan can be considered a K+-responsive smart polymer (Roy and Gupta, 2003). Mitsumata et al. (2003) have described the pH response of complex hydrogels made up of κ-carrageenan, chitosan, and CM-cellulose. The maximum degree of swelling was observed in the range of pH 11-12.

5.4.4

Applications

It was estimated that the worldwide sale of carrageenan in 2000 was around US$310 million (Van de Velde and Ruiter, 2002). The commercial applications of carrageenan revolve around their use as gelling, thickening and stabilizing agents. Processed food products such as ice creams, whipped cream, yogurt, jellies and sauces are some illustrative examples (Van de Velde et al., 2002). There are a number of reasons which make carrageenan an ideal component in food. Traditional use for > 600 years obviously initiated these applications in the industrial society. It is regarded as a GRAS item and has FDA (USA) approval for use as a food additive. In fact, the WHO expert committee recommended that it is not necessary to specify a daily limit for carrageenans (Van de Velde et al., 2002). However, a minimum average molecular weight of 100 kDa is prescribed since cecal and colonic ulceration was reported with fragments of carrageenan (Van de Velde and Ruiter, 2002). The carrageenan binds water nicely and this helps in formulations which have to contain aqueous fluids. Although not a surfactant, it does stabilize emulsions and suspensions. At high temperature, it melts and has lower viscosity. This allows processing and good heat transfer while dealing with food systems. Below 49°C, it solidifies and forms gels. The gels are stable at room temperature. Carrageenans have a better textural, mouthfeel, flavor, and processing properties as compared to starch. Thus, they can replace starch as thickener in many food preparations. In fact, κ-carrageenan increases the viscosity of starch systems manifold. Similarly, it shows synergism with locus bean gum and konjac flour and stronger elastic gels are obtained.

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An important property of carrageenan which makes this a better hydrocolloid to be used (in food and other systems) is the way it interacts with other proteins, especially caseins, the milk proteins. The positive charges on casein micelles interact electrostatically with the negatively charged sulfate of carrageenan and leads to stable and strong gels. Chocolate milk and flans are two examples of products which are based upon this interaction. Apart from food systems it is also being used in toothpastes and air fresheners. Hand lotions, shampoos and contraceptive gels represent growing/ potential market segments for carrageenans (http://www.micchem.com/products/ Carrageenan.htm). Carrageenans have been used for immobilization of whole cells and enzymes (Van de Velde et al., 2002). As the enzymes, in general, can leak out through porous carrageenan gels, gel hardening by use of K+ (high concentration), Ca2+, Al3+, Fe2+, galactomannans or glucomannans is required for this application. Van de Velde et al. (2002) have listed the applications of enzymes immobilized in carrageenan gels for various biotransformations. In addition, some bioanalytical applications, in H2O2 determination (immobilized catalase), pesticide analysis (co-immobilized choline oxidase and choline esterase), lecithin analysis in food and drugs (co-immobilized choline oxidase and choline esterase) and monitoring the rancidity of olive oils (immobilized tyrosine), have been mentioned (Van de Velde et al., 2002). Among the applications of whole cell immobilization in κ-carrageenan are waste water treatment, asymmetric synthesis and production of vinegar, milk prefermentation and production of beer and ethanol (Van de Velde et al., 2002). Applications of carrageenan as an excipient in drug formulations and other medical applications have been covered by Van de Velde and Ruiter (2002). Thommes et al. (2007) have recently examined the effect of drying on extruded pellets in which κ-carrageenan was used as a pelletization aid. It was found that heating above 80°C decreased the disintegration time. This has implications in the context of the drug release properties of κ-carrageenan pellets. More importantly, these authors suspect fragmentation of κ-carrageenan. In view of carrageenan fragments being not acceptable by WHO (as already mentioned), this is of serious concern and needs to be investigated carefully. The smart nature of κ-carrageenan as a polymer, has been exploited for bioseparation of pullulanase (Roy and Gupta, 2003) and yeast alcohol dehydrogenase (Mondal et al., 2003). In both cases, affinity precipitation (Roy and Gupta, 2002) was used as the bioseparation technique. The precipitation of κ-carrageenan was carried out by K+ addition. While the polysaccharide as such showed the selective affinity for pullulanase, in the other case, κ-carrageenan was used as a smart carrier for the dye cibacron blue which functioned as an affinity ligand. As other dyes in particular and affinity ligands in general can be linked to κ-carrageenan in a similar way, the strategy can be extended for purifying other enzymes as well.

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Other miscellaneous smart polysaccharides and their applications

While the three polysaccharides discussed previously have been more extensively used, some other polysaccharides (though used less frequently) have also been used for some interesting applications. Colon-specific drug release systems exploit the change in the pH along the gastrointestinal tract between 2 (stomach) and 10 (colon). Aguilar et al. (2007) mention the use of several polysaccharides (amylose, guar gum, pectin, inulin, chondroitin sulphate, dextran and locust bean gum) in designing colon-specific drug release systems. Zhang et al. (2007) have recently described a dextran based antigen/ antibody hydrogel. The presence of free antigen affected the antigen/antibody internal interactions and resulted in changes in the permeability of solutes through the membrane. However, it may be noted that it is not dextran from which the smart behavior originated, it was the smartness of the well known biological affinity pair of antigen/antibody. Themoreversibility of xyloglucan gels have been exploited in quite a few cases for obtaining drug release systems. This polysaccharide is obtained from tamarind seeds. It consists of a [1→4]-β-D-glucan backbone with [1→6]-α-D-xylose branches partially substituted by [1→2]-β-D-galactoxylose. Treatment of naturally occurring xyloglucan by β-galactosidase gives a thermally reversible xyloglucan gel whose sol-gel temperature can be varied by varying degree of hydrolysis. It is believed that xyloglycan gels may be useful for rectal and intraperitoneal drug delivery. Their usefulness in oral drug delivery has also been explored (Kumar et al., 2002). Gellan gum (produced by Pseudomonas elodea) is a linear anionic polymer of a repeating tetrasaccharide unit of glucose (two units), glucuronic acid and rhamnose. In the native state, some of the glucoses are acylated with acetyl and L-glyceryl groups. The viscosity of gellan gum dispersions is dependent upon pH, temperature, and the presence of cations. The gum is used in the food industry, as a growth media for bacteria and in plant tissue culture. Again, it has been used for designing sustained release systems for drugs (Kumar et al., 2002). Vigo (1998) has viewed the variety of structures which could be created by interacting cellulose with other polymers. A recent work shows that methylcellulose is an effective thermosensitive flocculant (Franks, 2005). Zohuriaan-Mehr et al. (2006) have described a hybrid hydrogel of gum arabic-acrylate which showed swelling-deswelling response to pH, salinity, Ca2+ and organic solvents. Garner et al. (1999) have described a polypyrrole-heparin composite in which exposed heparin could be varied by either application of negative potentials or by exposure to an aqueous reductant. While there is no ambiguity about what constitutes a smart material, it is

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sometime possible to confer smartness on a seemingly nonsmart material. For example, a cyclodextrin microgel was found to show a pH-dependent host-guest inclusion effect for a dye (Liu et al., 2004). Considering that cyclodextrins are already exploited extensively, this creates another dimension which will further their usefulness in many areas.

5.6

Polysaccharide-based composite materials

The previous discussion has focused on polysaccharide materials. In material science, it is not uncommon to blend, complex, copolymerize different materials to improve upon the desirable property of a polymeric material. In the area of smart materials also, many composite materials have been obtained from different polymeric materials. This section focuses on such materials wherein at least one component is a polysaccharide.

5.6.1

Examples

Many stimuli-responsive hydrogels have been prepared by combining polysaccharides (chitosan, alginate, cellulose and dextran) with thermoresponsive materials. The areas of potential application for such composite materials include drug delivery, tissue engineering and wound healing. In some cases where both components are smart, the composites show dual stimuli-responsive behavior. The preparative strategies used for obtaining such composite materials include graft copolymerization, blending, formation of polyelectrolyte complexes and core-shell type polymerization. A recent review (Prabaharan and Mano, 2006) deals with these approaches quite well and describes some of the composite materials. A non-toxic and biocompatible material was obtained by grafting poly(Nisopropyl acrylamide (NIPAAm)) monomer onto chitosan using ceric ammonium nitrate as the initiator. The copolymer had a lower critical solution temperature (LCST) of 32°C with a swelling ratio higher at pH 4 than at pH 7. The pH dependent swelling behavior was more noticeable at 25°C than at 32°C (Chung et al., 2005). Also, chitosan-g-pNIPAAm particles prepared by emulsion copolymerization have been reported. Again, the particles displayed dual stimuli-response as far as swelling behavior was concerned. A semiinterpenetrated network was obtained by the free radical polymerization of NIPAAm in the presence of chitosan by using tetraethyleneglycoldiacrylate as the crosslinking agent. The resulting material showed a dramatic response (in terms of degree of swelling) to pH. Response behaviors of such a semiinterpenetrated network and corresponding full-interpenetrated network have been found to be very different (Verestiue et al., 2004). A limiting factor for the use of pNIPAAm hydrogels in their applications in the areas of sensors, actuators and chemical valves has been the slow

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deswelling rate of pNIPAAm hydrogels. Semi-IPN hydrogels prepared from linear alginate and cross-linked pNIPAAm have shown better response rates. The cellulose-reinforced hydrogels showed the interesting property of pore size control with temperature. The pNIPAAm grafted to dextran formed micelles (spheres with mean diameter of < 30 nm) in aqueous media which, in principle, can be used for drug delivery of lipophilic drugs. Graft copolymers combining mostly pNIPAAm with other polysaccharides have also been prepared by radiation (e.g. UV, γ-irradiation) based methods and condensation reactions. Worth mentioning are the comb-type graft hydrogels obtained from alginate and pNIPAAm. These macroporous hydrogels showed rapid swelling/deswelling response to changes in both pH and temperature. Physically or chemically cross-linked polymeric blend based hydrogels have also been described. An interesting example is that of porous hydrogels with cell attachment and growth sites. The IPN hydrogels prepared from Ca2+-alginate and pNIPAAm showed different pore morphologies depending upon the temperature. The porous hydrogels became nonporous beyond their LCST temperature. The mechanical strength of these hydrogels also increased dramatically in their more compact form beyond their LCST. While glutaraldehyde has been more frequently used for obtaining chemically crosslinked blends (such as chitosan/pNIPAAm blends), genipin has also been used as a crosslinker for obtaining a blend of chitosan and poly(vinyl pyrrolidine) (PVP). Low pH and high temperature led to greater swelling which was also enhanced as PVP content was increased (Khurma et al., 2005). Dual sensitive polyelectrolyte complexes (PEC) have been prepared by combining cationic chitosan and anionic alginate with polymers carrying opposite charges. PECs are reported to have applications as membranes, antistatic coatings, sensors and medical prosthetic materials (Etienne et al., 2005; Casalbore-Miceli et al., 2006; Vasiliu et al., 2005). Core-shell type copolymers constitute a highly complex design in composite materials with the attractive property that the responsiveness is tunable (Prabaharan and Mano, 2006). Smart microgels with a thermoresponsive core with pH sensitive shells made up of pNIPAAm and chitosan have been described. Similarly, composites with a cross-linked copolymer of NIPAAm and chitosan as core and acrylate copolymers as shells have also been described. The main focus of studies on such microgels has been, of course, on studying their swelling/deswelling behavior at various pH values. The main application of these composite materials has been to design drug delivery systems which release drugs in a controlled fashion at a specific site. The biocompatible nature of chitosan and alginate has resulted in their being components of many composites synthesized for this application. A PEG-g-chitosan preparation which was an injectable liquid at low temperature but turned into a semisolid gel at body temperature showed linear release of BSA up to 70 hours (Bhattarai et al., 2005). A thermosensitive hydrogel

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made up of chitosan and β-glycerophosphate was found to work well as a site directed, injectable and controlled-release (over a 1 month period) system in a preclinical trial for paclitaxal delivery to localized solid tumors (Prabaharan and Mano, 2006). A number of studies have been reported with composites of alginate and pNIPAAm. In addition to response to the presence of Ca2+, the effect of pH on ionization of carboxyl groups of alginate was also exploited in such designs. The temperature determined the drug release rate and this dependence on the temperature itself could be varied by changing the pH. Some of the other thermoresponsive composites which showed promising results as drug delivery materials are ethylcellulose/pNIPAAm microcapsules and NIPAAm grafted on dextran methacrylate (Ichikawa and Fukumori, 2000; Huang and Lowe, 2005). One of the challenges in tissue engineering is to find a material which could serve as a cell culture carrier and allow harvesting in the case of highly adhesive mesenchymal stem cells. Chitosan-g-pNIPAAm has turned out to be a useful material; the cells could be harvested simply by lowering the temperature. The injectable composite material served as a good scaffold for chondrogenic differentiation of human stem cells (Prabaharan and Mano, 2006). Non-woven fabrics made of thermosensitive composites based upon chitosan also show promise as wound healing dressing materials. Such materials, interestingly, showed higher bacteriostatic property as compared to chitosan (Prabaharan and Mano, 2006). Various starch-based composites have been described in the literature (Marques et al., 2002). An extensive list of work on starch-based composites by the group of Prof. Reis can be accessed at http://www.3bs.uminho.pt. The target applications are in drug delivery (Malafaya et al., 2001) and tissue engineering (Gomes et al., 2001, Salgado et al., 2002). Such composite materials also include starch-chitosan hydrogels (Baran et al., 2004). The starch-based thermoplastic hydrogels used as bone cements and drug delivery carriers may also be mentioned here (Pereira et al., 1998). Finkenstadt (2005) has reviewed the applications of polysaccharides in designing biosensors, environmentally sensitive membranes and components in high-energy batteries. These applications are based upon their electroactive nature which is exploited by using them for doping, blending or grafting into other materials. While direct exploitation of smart behavior is not yet seen, it may turn out to be an asset. For example, electroactive polypyrroles required a negatively charged counterion hyaluronic acid to exhibit full conductivity. The composite laminate showed sharper responses in terms of cell compatibility, nontoxicity and increased vascularization as compared to the material without hyaluronic acid. Considering the intended application for tissue engineering, it may be possible to build-in specific cell responses toward stimuli.

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Cascone and Maltinti (1999) have evaluated blends of poly(vinyl alcohol) (PVA) with chitosan or dextran as drug delivery systems for growth hormone. The hormone helps in wound healing and tissue repair. The blended hydrogels were superior to pure PVA hydrogels with respect to release of PVA as such over a period of time. They are also less expensive than similar blends of PVA with collagen and hyaluronic acid which have been described earlier (Giusti et al., 1993; Guerra et al., 1994). It was found that either chitosan or dextran content controlled the hormone release. The dextran containing hydrogels reached the swelling equilibrium faster than chitosan containing blends. This has implication for drug release kinetics (Cascone and Maltinti, 1999). The initial step of water uptake is followed by the drug release step. Hence, the GH release clearly shows two-step kinetics in the case of chitosanPVA hydrogel whereas GH release appears as a single phase process for dextran-PVA hydrogel (Cascone and Maltinti, 1999). The same group, more recently, evaluated hydrogel blends of PVA with hyaluronic acid, dextran and gelatin as potential tissue engineering scaffolds (Cascone et al., 2004). PVA has been a material of choice as PVA hydrogels have water contents similar to those of natural tissues. Besides, PVA hydrogels are biocompatible, sterilizable and easy to mould into a desired shape. The blending was aimed at improving mechanical stability and biocompatibility. It was found that hydrogels containing dextran/ PVA in the ratio of 40:60 showed the highest porosity among all the blends tested. Overall, the blends showed the desired porosity for fibroblast growth. Whether these porosities will actually translate into support for cell adhesion and proliferation needs to be tested. Zhang et al. (2004) have synthesized dextran-maleic anhydride/pNIPAAm smart hybrid gels by photocrosslinking. FT-IR, DSC, swelling kinetics showed that the composite hydrogels were responsive to both temperature and pH. The LCST could be adjusted by changing the ratio of the two components during synthesis. Blending with dextran made these composite gels partially biodegradable. Finally, the work of Kaffashi et al. (2005), while preliminary in nature, illustrates the possibility of blending naturally occurring gums with more well defined polymer. These workers blended gelatin with tragacanth gum. This gum, isolated from the Astragalus plant consists of polygalacturonic acids. The smartness of the hydrogel was not investigated but at a conceptual level, this raises several interesting possibilities as a variety of plant gums have been described in the literature (Aspinall, 1969; Verbeken et al., 2003). Some more examples of composite materials based upon polysaccharides are shown in Table 5.3.

5.7

Future trends

Currently, materials based upon smart polysaccharides are extensively used in the food industry and to a lesser extent in some other industries. In

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Table 5.3 Some more examples of composite materials based upon polysaccharides Composite

Application

Reference

Chitin and chitosan based materials Poly-l-lysine coated covalently on chitosan beads

Adsorption of bilirubin

Chandy and Sharma (1992)

DNA-chitosan complexes

Removal/ concentration of carcinogenic heterocyclic amines

Hayatsu et al. (1997)

Chitosan conjugated magnetite

Recovery of recombinant E. coli

Honda et al. (1999)

Chitosan-sialic acid branched polysaccharides

Soluble hybrids Bound lectins

Sashiwa et al. (2000)

Chitosan-magnetite aggregates containing Nitrosomonas europaea cells

Ammonia removal from waste water

Liu et al. (2000)

Chitosan-hydroxyapatite composites

Bone substitute as bioceramics

Finisie et al. (2001)

Chitosan attached to sugar, dendrimers, cyclodextrins, crown ethers

Miscellaneous applications Sashiwa and Aiba including drug delivery (2004) systems and other medical applications

Alginate-chitosan-poly (lactic co-glycolic acid) composite microspheres

Protein delivery systems

Zheng et al. (2004)

Nanostructured poly (lactic-co-glycolic acid)/chitin matrix

Tissue engineering

Min et al. (2004)

Self-assembled Immobilized chitosan/poly organophosphorus (thiophene-3-acetic acid) hydrolase for detection layers of paraoxon

Alginates Dried calcium alginate/ magnetite spheres

Constantine et al. (2003)

Support for chromatographic Burns et al. (1985) separations and enzyme immobilization

A mixed gel of colloidal silica and alginate

Ethanol production by yeast immobilized in the mixed gel

Fukushima et al. (1988)

Chitosan-alginate coacervate capsules

Encapsulation of cells/ tissues/pharmaceuticals

Daly and Knorr (1988)

Alginate-starch copolymers

Affinity adsorption of α-amylase

Somers et al. (1993)

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Table 5.3 (Continued) Composite

Application

Reference

Xanthan-alginate spheres

Encapsulation of urease

Elcin (1995)

Polyethyleneiminemodified barium alginate

Immobilization of cephalosporium acremonium for production of cephalosporin C

Park and Khang (1995)

Alginate beads coated with chitosan or DEAE-dextran

Protein release system

Huguet et al. (1996)

Alginate-polythylene glycol gels

Cultivation of mammalian cells

Seifert and Phillips (1997)

Alginate-polylysine capsules

Immunoprotection of endocrine cells

De Vos et al. (1997)

Poly(methylene co-guanidine) coated alginate

Encapsulation of urease

Hearn and Neufeld (2000)

Alginate-chitosan beads

Immobilization of antibodies

Albarghouthi et al. (2000)

Alginate-Konjac glucomannanchitosan beads

Controlled release system for proteins

Wang and He (2002)

Multilayer alginate/ protamine microsized capsules

Encapsulation of α-chymotrypsin

Tiourina and Sukhorukov (2002)

Magnetized alginate

Magnetic resonance imaging

Shen et al. (2003)

Magnetic alginate particles Purification of α-amylase

Safarikova et al. (2003)

Alginate-chitosan coreshell microcapsules

Enzyme immobilization

Taqieddin and Amiji (2004)

Additive for low fat beef frankfurters

Candogen and Kolasarici (2003)

κ-Carrageenan-g-poly acrylamide

Adsorption of fluids and adhesion

Meena et al. (2006)

Carrageenan-g-poly (Sodium acrylate)/ kaolin hydrogels

Superabsorbent composites

Pourjavadi et al. (2007)

Carrageenans Carrageenan-pectin gels

other areas like biosensors, molecular gates and valves, the synthetic thermostable polymer pNIPAAm has dominated. Increasingly, the composites of synthetic polymers and polysaccharides are being investigated for their applications as well as in designing drug release systems. In tissue engineering

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and other usages wherein biocompatibility is a key factor, polysaccharides scores over synthetic smart polymers. Again, composites may be the ideal materials. The information given in this chapter hopefully will motivate research workers to more vigorously exploit polysaccharides in designing smart materials. There are two more reasons to use polysaccharides more often. The current realization that marine biodiversity offers a rich source of materials should lead to a search for a near ideal polysaccharide for a particular purpose. Nature already had made those ‘combinatorial libraries’ of diverse structures. Second, in the drive towards a sustainable society, biodegradable materials from renewable resources constitute an important class. Polysaccharides are from renewable sources and are biodegradable to a varying extent. A survey of the recent patented literature shows that a trend of using polysaccharides for niche applications is emerging. A recent US patent uses cellulose derivatives for forming an ink receptive top layer on materials used for recording inkjet images (Baker, 2003). Another US patent (Ni and Yates, 2004) uses sodium alginate to improve gelation properties of pectic substances for delivery of basic fibroblast growth factor. Some more examples can be found in a review by Al-Tahami and Singh (2007). Given rich structural biodiversity, easy possibility of conjugation/complexation of other substances, biodegradability and biocompatibility (to a varying degree depending upon the particular polysaccharide), polysaccharides and composites based upon polysaccharides are bound to find increasing numbers of applications in diverse areas. Their smartness in many cases is an additional attractive feature.

5.8

Acknowledgement

The preparation of this chapter and the research work from the authors’ laboratory mentioned in this chapter were supported by the Department of Science and Technology (Government of India) core group grant on ‘applied biocatalysis’ and Department of Biotechnology (Government of India) project grants. The support by the Indian Council of Medical Research in the form of Senior Research Fellowship to SR is also acknowledged.

5.9

Sources of further information

A search on Google Scholar™ beta with the phrase ‘Stimuli-sensitive polysaccharides’ yielded about 6530 hits. The sources varied from biotechnology journals to microbiology journals or medical journals. This reflects the wide range of relevance of this broad class of materials. It also conveys that this area has become truly an area which can immensely benefit from multidisciplinary efforts. Some of the sources which we would like to recommend are:

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General references on smart materials Roy I and Gupta M N (2003), ‘Smart polymeric materials: Emerging biochemical applications’, Chem Biol, 10, 1161–1171. Hoffman A S (2002), ‘Hydrogels for biomedical applications’, Adv Drug Deliv Rev, 43, 1–12. Peppas N A (1985), Hydrogels in medicine and pharmacy, Boca Raton, FL, CRC Press.

General references on bioseparation by using smart polysaccharides Gupta M N (ed.) (2002), Methods in Affinity-based Separation of Proteins/ enzymes, Switzerland, Birkhauser Verlag. Mondal K, Roy I and Gupta M N (2006), ‘Affinity based strategies for protein purification’, Anal Chem, 78, 3499–3504. Roy I, Mondal K and Gupta M N (2007), ‘Leveraging protein purification strategies in proteomics’, J Chromatogr B, 849, 32–42. Mondal K and Gupta M N (2006), ‘The affinity concept in bioseparation: Evolving paradigms and expanding range of applications’, Biomol Eng, 23, 59–76.

Chitosan and chitin Kumar M N V R (1999), ‘Chitin and chitosan fibres: A review’, Bull Mater Sci, 22, 905–915. Shahidi F, Kamil J, Arachchi V and Jeon Y J (1999), ‘Food applications of chitin and chitosans’, Trends Food Sci Technol, 10, 37–51. Muzzarelli R A A (1977), Chitin, Oxford, Pergamon Press. Skjåk-Bræk G, Anthonsen T and Sandford P (eds) (1989), Chitin and Chitosan, London, Elsevier. http://wwwcsi.unian.it/chimicam/chimicam.html

Alginates Gerbsch N and Buchholz R (1995), ‘New processes and actual trends in biotechnology’, FEMS Microbiol Rev, 16, 259–269. (A good and informative review of immobilization techniques with special emphasis on alginate.) Martinsen A, Skjåk-Bræk G and Smidsrød O (1989), ‘Alginate as immobilization material: I. Correlation between chemical and physical properties of alginate gel beads’, Biotechnol Bioeng, 33, 79–89.

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Skjåk-Bræk G, Murano E and Paoletti S (1989), ‘Alginate as immobilization material. II: Determination of polyphenol contaminants by fluorescence spectroscopy, and evaluation of methods for their removal’, Biotechnol Bioeng, 33, 90–94. Smidsrød O and Skjåk-Bræk G (1990), ‘Alginate as immobilization material for cells’, TIBTECH, 8, 71–78.

κ-carrageenans Van de Velde F and De Ruiter G A (2002), ‘Polysaccharides from eukaryotes’, in Biopolymers, Vol 6, Polysaccharides II, Weinheim, Wiley-VCH, 245– 274. Van de Velde F, Lourenço N D, Pinheiro H M and Bakker M (2002), ‘Carrageenan: A food-grade and biocompatible support for immobilisation techniques’, Adv Synth Catal, 344, 815–835. http://www.fmcbiopolymer.com/PopularProducts/FMCCarrageenan/ Introduction/tabid/804/Default.aspx.

Composites Kumar M N V R, Kumar N, Domb A J and Arora M (2002), ‘Pharmaceutical polymeric controlled drug delivery systems’, in Advances in Polymer Science, Vol 160, Heidelberg, Springer Verlag. Aguilar M R, Elvira C, Gallardo A, Vázquez B and Román J S (2007), ‘Smart polymers and their applications as biomaterials’, Topics in Tissue Engineering, 3, 1–27. Al-Tahami K and Singh J (2007), ‘Smart polymer based delivery systems for peptides and proteins’, Recent Pat Drug Del Formul, 1, 65–71.

5.10

References

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Downs E C, Robertson N E, Riss T L and Plunkett M L (1992), ‘Calcium alginate beads as a slow-release system for delivering angiogenic molecules in vivo and in vitro’, J Cell Physiol, 152(2), 422–429. Draget K I, Myhre S, Skjåk-Brøk G and Østgaard K (1988), ‘Regeneration, cultivation and differentiation of plant protoplasts immobilized in Ca-alginate beads’, J Plant Physiol 132, 552–556. Draget K I, Varum K M, Moen E, Gynnild H and Smidsrod O (1992), ‘Chitosan crosslinked with Mo(Vi) polyoxyanions – A new gelling system’, Biomaterials, 13, 635– 638. Dutta P M (ed.) (2005), Chitin and Chitosan, Opportunities and Challenges, India, Contai. Elcin Y M (1995), ‘Encapsulation of urease enzyme in xanthan-alginate spheres’, Biomaterials, 16(15), 1157–1161 Etienne O, Gasnier C, Taddei C, Voegel J C, Aunis D, Schaaf P, Metz-Boutigue M H and Bolcato-Bellemin A L (2005), ‘Antifungal coating by biofunctionalized polyelectrolyte multilayered film’, Biomaterials, 26, 6704–6712. Fathy M, Safwat S M, el-Shanawany S M, Tous S S and Otagiri M (1998), ‘Preparation and evaluation of beads made of different calcium alginate compositions for oral sustained release of tiaramide’, Pharm Dev Technol, 3(3), 355–364. Finisie M R, Josue A, Favere V T and Laranjeira M C (2001), ‘Synthesis of calciumphosphate and chitosan bioceramics for bone regeneration’, An Acad Bras Cienc, 4, 525–532 Finkenstadt V L (2005), ‘Natural polysaccharides as electroactive polymers’, Appl Microbiol Biotechnol, 67, 735–745. Franks G V (2005), ‘Stimulant sensitive flocculation and consolidation for improved solid/liquid separation’, J Colloid Interface Sci, 292, 598–603. Fukushima Y, Okamura K, Imai K and Motai H (1988), ‘A new immobilization technique of whole cells and enzymes with colloidal silica and alginate’, Biotechnol Bioeng, 32, 584–594. Garner B, Georgevich A, Hodgson A J, Liu L and Wallace G G (1999), ‘Polypyrroleheparin composites as stimulus-responsive substrates for endothelial cell growth’, J Biomed Mater Res, 44, 121–129. Giusti P, Lazzeri L, Barbani N, Narducci P, Bonaretti A, Palla M and Lelli L (1993), ‘Hydrogels of poly (vinyl alcohol) and collagen as new bioartificial materials’, J Mater Sci Mater Med, 4, 538–542. Gomes M E, Ribeiro A S, Malafaya P B, Reis R L and Cunha A M (2001), ‘A new approach based on injection moulding to produce biodegradable starch-based polymeric scaffolds: morphology, mechanical and degradation behaviour’, Biomaterials, 22, 883–889. Guisan J M (2006), Immobilization of enzymes and cells, Totowa, NJ, Humana Press. Gupta M N (1992), ‘Enzyme function in organic solvents’, Eur J Biochem, 203, 25– 32. Gupta M N and Mattiasson B (1994), ‘Affinity precipitation’, in Street G, Highly Selective Separations in Biotechnology, London, Chapman and Hall, 7–33. Gupta M N (ed.) (2000), Methods in Non-aqueous Enzymology, Switzerland, Birkhauser Verlag. Gupta M N (ed.) (2002), Methods in Affinity-based Separation of Proteins/enzymes, Switzerland, Birkhauser Verlag. Haque A and Morris E R (1993), ‘Thermogelation of methylcellulose, Part I: molecular structures and processes’, Carbohydr Polym, 22, 161–173.

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Hatti-Kaul R (2000), Aqueous Two Phase Systems – Methods and Protocols, Totowa, NJ, Humana Press. Hayatsu H, Tanaka Y and Negishi K (1997), ‘Preparation of DNA-chitosan columns and their applications: binding of carcinogens to the column’, Nucleic Acids Symp Ser, 37, 139–140. Hearn E and Neufeld R J (2000), ‘Poly (methylene co-guanidine) coated alginate as an encapsulation matrix for urease’, Proc Biochem, 35, 1253–1260. Henriksen I, Skaugrud Ø and Karlsen J (1993), ‘Use of chitosan and chitosan malate as an excipient in wet granulation of three water soluble drugs’, Int J Pharm, 98, 181– 188. Hoffman A S (2002), ‘Hydrogels for biomedical applications’, Adv Drug Deliv Rev, 43, 1–12. Honda H, Kawabe A, Shinkai M and Kobayashi T (1999), ‘Recovery of recombinant Escherichia coli by chitosan-conjugated magnetite’, Biochem Eng J, 3, 157– 160. Huang X and Lowe T L (2005), ‘Biodegradable thermoresponsive hydrogels for aqueous encapsulation and controlled release of hydrophilic model drugs’, Biomacromolecules, 6, 2131–2139. Huguet M L, Neufeld R J and Dellacherie E (1996), ‘Calcium-alginate beads coated with polycationic polymers: comparison of chitosan and DEAE-dextran’, Proc Biochem, 31(4), 347–353 Ichikawa H and Fukumori Y (2000), ‘A novel positively thermosensitive controlledrelease microcapsule with membrane of nano-sized poly (N-isopropylacrylamide) gel dispersed in ethylcellulose matrix’, J Controlled Release, 63, 107–119. Jain S, Mondal K and Gupta M N (2005), ‘Applications of alginate in bioseparation of proteins’, Artif Cell Blood Substit Biotechnol, 34, 127–144. Jeong B and Gutowska A (2002), ‘Lessons from nature: stimuli-responsive polymers and their biomedical applications’, TIBTECH, 20, 305–311. Kaffashi B, Khadiv-Parsi P and Zandieh A (2005), ‘Smart drug release implementing the Tragacanth and Collagen composite’, 8th Int Symp. Polymers for Advanced Technologies, Hungary. Kim J J and Park K (2002), ‘Applications of smart hydrogels in separation’, in Galaev I Y and Mattiasson B, Smart Polymers for Bioseparation and Bioprocessing, London, Taylor & Francis, 140–162. Khurma J R, Rohindra D R and Nand A V (2005), ‘Swelling and thermal characteristics of genipin crosslinked chitosan and poly(vinyl pyrrolidone) hydrogels’, Polym Bull, 54, 195–204. Krajewska B (2004), ‘Application of chitin- and chitosan-based materials for enzyme immobilizations: a review’, Enzyme Microb Tech, 35, 126–139. Kuera J (2004), ‘Fungal mycelium-the source of chitosan for chromatography’, J Chromatogr B, 808, 69–73. Kumar M N V R, Kumar N, Domb A J and Arora M (2002), ‘Pharmaceutical polymeric controlled drug delivery systems’, in Advances in Polymer Science, vol 160, Heidelberg, Springer Verlag. Liener I E, Sharon N and Goldstein I J (1986), The Lectins, New York, Academic Press, 437–466. Liu C, Honda H, Ohshima A, Shinkai M and Kobayashi T (2000), ‘Development of Chitosan-Magnetite Aggregates Containing Nitrosomonas europaea cells for Nitrification Enhancement’, J Biosci Bioeng, 89(5), 420–425

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Liu Y-Y, Fan X-D, Kang T and Sun L (2004), ‘A cyclodextrin microgel for controlled release driven by inclusion effects’, Macromol Rapid Commun, 25, 1912–1916. Malafaya P B, Elvira C, Gallardo A and Roman J S (2001), ‘Porous starch-based drug delivery systems processed by a microwave route’, J Biomater Sci Polymer Edn, 12, 1227–1241. Marques A P, Reis R L and Hunt J A (2002), ‘The biocompatibility of novel starch-based polymers and composites: in vitro studies’, Biomaterials, 23, 1471–1478. Martinsen A, Skjåk-Bræk G and Smidsrød O (1989), ‘Alginate as immobilization material: I. Correlation between chemical and physical properties of alginate gel beads’, Biotechnol Bioeng, 33, 79–89. Masteiková R, Chalupová Z and Šklubalová Z (2003), ‘Stimuli-sensitive hydrogels in controlled and sustained drug delivery’, Medicina, 39, 19–24. Meena R, Prasad K, Mehta G and Siddhanta A K (2006), ‘Synthesis of the copolymer hydrogel κ-carrageenan-graft-PAAm: Evaluation of its absorbent and adhesive properties’, J Appl Polym Sci, 102, 5144–5153. Michel G, Nyval-Collen P, Barbeyron T, Czjzek M and Helbert W (2006), ‘Bioconversion of red seaweed galactans: a focus on bacterial agarases and carrageenases’, Appl Microbiol Biotechnol, 71, 23–33. Min B M, You Y, Kim J M, Lee S J and Park W H (2004), ‘Formation of nanostructured poly (lactic-co- glycolic acid)/chitin matrix and its cellular response to normal human keratinocytes and fibroblasts’, Carbohydr Pol, 57(3), 285–292. Mitsumata T, Suemitsu Y, Fujii K, Fujii T, Taniguchi T and Koyama K (2003), ‘pHresponse of chitosan, κ-carrageenan, carboxymethyl cellulose sodium salt complex hydrogels’, Polymer, 44, 7103–7111. Mondal K and Gupta M N (2006), ‘The affinity concept in bioseparation: Evolving paradigms and expanding range of applications’, Biomol Eng, 23, 59–76. Mondal K, Bohidar H B, Roy R P and Gupta M N (2006), ‘Alginate-chaperoned facile refolding of Chromobacterium viscosum lipase’, Biochim Biophys Acta, 1764, 877– 886. Mondal K, Roy I and Gupta M N (2003), ‘κ-carrageenan as a carrier in affinity precipitation of yeast alcohol dehydrogenase’, Protein Exp Purif, 32, 151–160. Ni Y and Yates K M (2004), ‘In situ gel formation of pectin’, US Patent No. 20046777000. Park H J and Khang Y H (1995), ‘Production of cephalosporin C by immobilized Cephalosporium acremonium in polyethyleneimine-modified barium alginate’, Enzyme Microb Technol, 17, 408–412. Peppas N A (1995), ‘Hydrogels in medicine and pharmacy’, Boca Raton Fl, CRC Press. Pereira C S, Cunha A M, Reis R L, Vazquez B and Roman J S (1998), ‘New starch-based thermoplastic hydrogels for use as bone cements or drug-delivery carriers’, J Mater Sci: Mater Med, 9, 825–833. Pourjavadi A, Hosseinzadeh H, Mahdavinia G R and Zohuriaan-Mehr M J (2007), ‘Carrageenan-g-poly(sodium acrylate)/kaolin superabsorbent hydrogel composites: synthesis, characterization and swelling behaviour’, Polymers and Polymer Composites, 15, 43–51. Prabaharan M and Mano J F (2006), ‘Stimuli-responsive hydrogels based on polysaccharides incorporated with thermo-responsive polymers as novel biomaterials’, Macromol Biosci, 6, 991–1008. Roy I and Gupta M N (2002), ‘Macro-(afinity ligands) in bioseparations’ in Gupta M N, (ed.) Methods for Affinity-based Separation of Proteins/enzymes, Basel, Birkhauser Verlag, 130–147.

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Roy I and Gupta M N (2003), ‘κ-carrageenan as a new smart macro-(affinity ligand) for the purification of pullulanase’, J Chromatogr A, 998, 103–108. Roy I and Gupta M N (2003), ‘Smart polymeric materials: emerging biochemical applications’, Chem Biol, 10(12), 1161–1171. Roy I, Sharma S and Gupta M N (2004), ‘Smart biocatalysts: design and applications’, Adv Biochem Engin/Biotechnol, 86, 159–189. Šafařıˇkova M, Roy I, Gupta M N and Šafařıˇk I (2003), ‘Magnetic alginate microparticles for purification of α-amylases’, J Biotechnol, 105, 255–260. Salgado A J, Gomes M E, Chou A, Coutinho O P, Reis R L and Hutmacher D W (2002), ‘Preliminary study on the adhesion and proliferation of human osteoblasts on starchbased scaffolds’, Mater Sci Engin C, 20, 27–33. Sashiwa H and Aiba S I (2004), ‘Chemically modified chitin and chitosan as biomaterials’, Prog Polym Sci, 29, 887–908. Sashiwa H, Makimura Y, Shigemasa and Roy R (2000), ‘Chemical modification of chitosansialic acid branched polysaccharide hybrids’, Chem Commun, 909–910. Seifert D B and Phillips J A (1997), ‘Porous alginate-poly (ethylene glycol) entrapment system for the cultivation of mammalian cells’, Biotechnol Prog, 13(5), 569–576. Senstad C and Mattiasson B (1989), ‘Purification of wheat germ agglutinin using affinity flocculation with chitosan and a subsequent centrifugation or flotation step’, Biotech Bioeng, 34, 387–393. Shahidi F, Kamil J, Arachchi V and Jeon Y J (1999), ‘Food applications of chitin and chitosans’ Trends Food Sci Technol, 10, 37–51. Sharbati Del, Guerra R, Cascone M G, Barbani N and Lazzeri L (1994), ‘Biological characterization of hydrogels of poly(vinyl alcohol) and hyaluronic acid’, J Mater Sci Mater Med, 5, 613–616. Shen F, Legrand C P, Somers S, Slade A, Yip C, Duft A M, Winnik F M and Chang P L (2003), ‘Properties of a novel magnetized alginate for magnetic resonance imaging’, Biotechnol Bioeng, 83(3), 282–292. Singh D K and Ray A R (2000), ‘Biomedical applications of chitin, chitosan and their derivatives’, Rev Macromol Chem Phys, 40, 69–83. Skjåk-Bræk G, Murano E and Paoletti S (1989), ‘Alginate as immobilization material. II: Determination of polyphenol contaminants by fluorescence spectroscopy, and evaluation of methods for their removal’, Biotechnol Bioeng, 33, 90–94. Smidsrød O and Draget K I (1997), ‘Alginate gelation technologies’, in Dickinson E and Bergenståhl B, Food Colloids, Cambridge, Royal Society of Chemistry, 279– 293. Smidsrød O and Skjåk-Bræk G (1990), ‘Alginate as immobilization material for cells’, TIBTECH, 8, 71–78. Somers W A C, Lojenga A K, Bonte A, Rozie H J, Visser J, Rombouts F M and Riet K V (1993), ‘Alginate-starch co-polymers and immobilized starch as affinity adsorbents for α-amylase’, Biotechnol Appl Biochem, 18, 9–24. Taqieddin E and Amiji M (2004), ‘Enzyme immobilization in novel alginate-chitosan core-shell microcapsules’, Biomaterials, 25, 1937–1945. Teotia S and Gupta M N (2001a), ‘Free polymeric bioligands in aqueous two phase affinity extractions of microbial xylanases and pullulanase’, Protein Exp Purif, 22, 484–488. Teotia S and Gupta M N (2001b), ‘Reversibly soluble macro-(affinity ligand) in aqueous two phase separation of enzymes’, J Chromatogr A, 923, 275–280.

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Terbojevich M, Muzzarelli R A A (2000), ‘Chitosan’, in Phillips G and Williams P, Handbook of Hydrocolloids, Cambridge, Woodhead, 367–378. Thommes M, Blaschek W and Kleinebudde P (2007), ‘Effect of drying on extruded pellets based on κ-carrageenan’, European J Pharm Sci, 31, 112–118. Tiourina O P and Sukhorukov G B (2002), ‘Multilayer alginate/protamine microsized capsules: encapsulation of α-chymotrypsin and controlled release study’, Int J Pharm, 242, 155–161. Tyagi R, Kumar A, Sardar M, Kumar S and Gupta M N (1996), ‘Chitosan as an affinity macroligand for precipitation of N-acetyl glucosamine binding proteins/ enzymes’, Isol Purif, 2, 217–226. Urry D W (1993), ‘Molecular machines: how motion and other functions of living organisms can result from reversible chemical changes’, Angew Chem, 22, 819–841. Van Damme J M, Peumans W J, Pusztai A and Bardocz S (1997), Handbook of plant lectins: properties and biomedical applications, Weinheim, John Wiley & Sons. Van de Velde F and De Ruiter G A (2002), ‘Polysaccharides from eukaryotes’, in A Steinbüchel, S De Baets and E J Vandamme (eds) Biopolymers, Vol 6, Polysaccharides II, Weinheim, Wiley-VCH, 245–274. Van de Velde F, Lourenço, N D, Pinheiro H M and Bakker M (2002), ‘Carrageenan: A food-grade and biocompatible support for immobilisation techniques’, Adv Synth Catal, 344, 815–835. Vasiliu S, Popa M and Rinaudo M (2005), ‘Polyelectrolyte capsules made of two biocompatible natural polymers’, Eur Polym J, 41, 923–932. Verbeken D, Dierckx S and Dewettinck K (2003), ‘Exudate gums: occurrence, production, and applications’, Appl Microbiol Biotechnol, 63, 10–21. Verestiue L, Ivanov C, Barbu, E and Tsibouklis J (2004), ‘Dual-stimuli-responsive hydrogels based on poly (N-isopropylacrylamide)/chitosan semi-interpenetrating networks’ Int J Pharm, 269, 185–194. Vigo T L (1998), ‘Interaction of cellulose with other polymers: retrospective and prospective’, Polymers Adv Technol, 9, 539–548. Vincent J F V (2000), ‘Smart by name, smart by nature’, Smart Mater Struct, 9, 255– 259. Vos P D, Haan B D, Schilfgaarde R V (1997), ‘Effect of the alginate composition on the biocompatibility of alginate-polylysine microcapsules’, Biomaterials, 18(3), 273– 278. Walter H and Johansson G (1994), Aqueous Two-phase Systems: Methods in Enzymology, Volume 228, New York, Academic Press. Wang K and He Z (2002), ‘Alginate-konjac glucomannan-chitosan beads as controlled release matrix’, Int J Pharm, 244, 117–126. Wang L, Shelton R M, Cooper P R, Lawson M, Triffit J T and Barralet J E (2003), ‘Evaluation of sodium alginate for bone marrow cell tissue engineering’, Biomaterials, 24, 3475–3481. Yang J, Goto M, Ise H, Cho C-S and Akaike T (2002), ‘Galactosylated alginate as a scaffold for hepatocytes entrapment’, Biomaterials, 23, 471–479. Zhang R, Bowyer A, Eisenthal R and Hubble J (2007), ‘A smart membrane based on an antigen-responsive hydrogel’, Biotech Bioeng, 97, 976–984. Zhang X, Wu D and Chu C-C (2004), ‘Synthesis and characterization of partially biodegradable, temperature and pH sensitive Dex-MA/PNIPAAm hydrogels’, Biomaterials, 25, 4719–4730.

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Part II Surface modification and biomimetic coatings

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6 Surface modification for natural-based biomedical polymers I. P A S H K U L E VA, P. M. L Ó P E Z - P É R E Z and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

6.1

Introduction

Surface is defined as the outside or top layer of the material. If the analogy with a human is used, one can say that the bulk properties of a material determine its ‘character’, while the surface is its ‘face’. Similar to human society, the initial acceptance or rejection of a biomaterial in the cell society is very dependent on its face whereas material character determines its long performance and proper function. Unfortunately, it is very difficult to find a biomaterial which simultaneously possesses both suitable mechanical properties in order to function properly in a certain bioenvironment and to not be harmful for the host tissue.1 Therefore, a common approach is to fabricate biomaterials with adequate bulk properties and then to make-up those by a specific treatment resulting in enhanced surface properties. The materials’ surfaces (as people’s faces) are very different and it is not possible to have a universal modification for all of them.2 Moreover, the environment and the role which a certain biomaterial is expected to play, call for a specific, unique and resistant enough modification to ensure its good performance. To make this task even more complex, the requirements in the biomedical material research and development field are growing very fast. While a few years ago, bioinert surfaces, protecting biomaterials from bacterial invasion, were sufficient for a material to be successful,3 over the past decade the requirements have shifted4 to surfaces that interact and functionally integrate with their biological environment in a predictable and controllable way. Nowadays, design of surfaces helping the body to heal itself5 by stimulating specific cellular responses at the molecular level is the target of the research.

6.2

Some terms and classifications

A crucial concept to understand about the tissue–biomaterial interface is that many things happen there! The environment inside the body is dynamic and 165 © 2008, Woodhead Publishing Limited

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active, and the interface between an implanted biomaterial and the body is the location of a variety of dynamic biochemical processes and reactions.6 During contact of non-bio surfaces with biological fluids, protein adsorption occurs almost instantaneously. This protein layer will further mediate the key bio/material interactions. Therefore, protein adsorption plays a fundamental role in dictating the cellular response elicited by biomedical systems implanted in the human body. Thus, the ability to control these phenomena at the biomaterial surface largely determines the biological performance of biomedical systems. Prevention of non-specific adhesion of proteins and polymer functionalization with cell-type specific molecules can help to direct control of cell adhesion on biomaterials.7 In the field of biomaterials, two historical approaches have been utilized8 to understand and tailor cell adhesion to the materials’ surfaces. The older one, so-called material approach, correlates cell response (morphology, adhesion, retention or higher cellular function) to the character of the material surface. Different chemistry and physics based methodologies have been developed (Table 6.1) in order to tailor material surface in terms of composition, surface energy, morphology, and chemistry. According to the second, biology-driven approach, cell/biomaterial surface interactions are governed by the same biologically specific chemistry as cell/cell surface interactions. Following this approach, the material surface must be designed in a way to mimic the cell surface as close as possible (Table 6.2). Intensive exploration/exploitation of cell surface and its different components (e.g. proteins, phospholipids, enzymes, etc.) was the outcome from the development of this approach. Nowadays, these two approaches have merged and combined methodologies using the best achievements from Table 6.1 Material approach: Some of the methods used and related references Process

Methods

References

Etching

Chemical Physical

[9, 15, 16, 86, 87] [30–32, 36, 88, 89]

Functionalization Oxidation

Chemical Physical • Plasma [1, 21, 27–29, 40, 41] • UV irradiation [40, 41]

[9–12] [30, 32, 34–38, 65, 88, 90] [41, 42, 56]

Hydrolysis

Chemical Enzymatic

[23–26] [91]

Coatings

LbL

[92, 93]

Grafting[47, 48]

Chemical Enzymatic Physical activation: • Plasma

[49–53, 64, 94–104] [105–107]

•

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Irradiation (gamma, UV, laser)

[31, 33, 34, 51, 64, 72, 108–110] [54–56, 86, 111–114]

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Table 6.2 Some bio-approaches: Methods and applications Targeted Application

Methods

References

Cell adhesion

Protein immobilization Active peptide sequences conjugation [69, 118]

[53, 64, 115–117] [96, 98, 119–121]

Drug delivery

Self assembled structures (e.g. phospholipid cell membranes) Other chemical approaches

[83, 84, 122–125] [126, 127]

biology and material sciences are used for directing the interaction between tissue cells and biomaterials.

6.3

Wet chemistry in surface modification

Chronologically, this is the first surface modification approach used in order to improve surface properties of polymers. The wet chemical methods in the surface modification field can be compared with cosmetic surgery (but not with a simple make up!!!) if the analogy material surface/human face is used. The ultimate goal of this approach is to create stable, well-defined functional substrates characterized by controlled surface properties, which are available for further chemistry. The wet chemistry surface modification methods are based on the knowledge from general solution chemistry. Thus, for example starch-based blends have been surface oxidized by the well known oxidizing system acid-permanganate9 or surface crosslinked using tri-sodium tri-meta phosphate solution;10 chitosan can be surface sulfonated by SO311 or surface phosphonated by P2O512 in different solvents. Although the experience from the solution chemistry is indispensable, several specific ‘surface issues’ must be considered: (a) Which are the functional groups available on the surface? Are they the same as the ones in the bulk? Surface chemistry depends on the processing of the material. Therefore, prior to any further modification, full surface characterization and knowledge of the processing ‘history’ of the material are required. When a solvent is involved in the preparation of the sample (e.g. solvent casting technique), the ability of the solvent used to form hydrogen bonds with the functional groups of the material can show up or hide these functional groups. Usually polar, protic solvents result in more hydrophilic surfaces compared to aprotic ones. On the other hand, the mould’s surface, which is in contact with the sample, has a similar effect via hydrophobic/hydrophilic forces. A simple example is the contact angle of PCL membranes prepared by solvent casting

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using different solvents: CHCl3 85.08 (Petri dish contact surface)/93.3 (air contact surface); THF 105.8 (Petri dish contact surface)/101.7 (air contact surface).13 (b) Where does the reaction actually occur? Are the wet chemistry methods surface modification methods? The dynamics of the surface chemical composition in the wet surface chemistry methods additionally complicates the process. In this case, a solvent is also involved in the modification step. Once again, its interactions with the material to be modified can alter the surface chemistry. Moreover, if these interactions result in swelling, the modification will not be confined to the surface and will go deeply into the bulk of the material. All these issues must be considered in the choice of a system/method for surface modification of a certain material. The most common wet modification methods and some general trends in their application are described below. It must be noted that these methods are widely used in industry to treat large objects that would be difficult to treat by other commonly used techniques.

6.3.1

Wet chemical etching

Etching is a process of removal of surface material, similar to face lifting. It has a long history, starting at the beginning of the Middle Ages. The old masters such as Rembrandt and Goya used it as one of the main techniques to create their art works. However, the ‘art application’ of this method is constricted only to metals as materials. The widely used micro- and nanofabrication techniques14 are based on the same principles. Natural-based polymers are much more sensitive and the strong acids usually used for etching metals or glass, cannot be applied to them. Generally, weaker chemical etchants such as diluted bases and acids,15 oxidizing agents9,16 are used to convert smooth hydrophobic surfaces to rough hydrophilic surfaces, usually by dissolution of amorphous regions and surface oxidation and hydrolysis. The alternative plasma etching or so-called dry etching is preferable for surface modification and surface cleaning of biopolymers.

6.3.2

Oxidation by wet surface modification methods

What is the role of the oxygen in the surface chemistry of the applied biomaterials? Do we want it there, on the surface, or not? Which is its optimal surface content? Usually the surface oxidation alters the proteins’ adsorption and therefore cell behaviour via: (a) Modulation of the surface hydrophilicity, i.e. the physical bonds surface/ proteins. Generally, the introduction of oxygen containing groups, such © 2008, Woodhead Publishing Limited

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as hydroxyl (–OH), carbonyl (–C=O) or carboxyl (–COOH) groups, is related to an increase of the surface’s hydrophilicity. (b) Alteration of the surface charge. Negatively charged groups have shown17–20 a good effect on cell adhesion and growth and this is attributed to the favourable protein conformation on these surfaces. The polarity of these groups allows formation of additional hydrogen bonds with the proteins, which will keep them fixed onto the surface. (c) Creating active places where a chemical bond between the proteins and surface functional groups can occur. However, this process is not always advantageous since denaturation of the proteins could also occur. As mentioned before, general knowledge from organic solution chemistry can be used and solutions with known oxidative properties can be adjusted (concentration) and applied. An example is the oxidation of starch-based biomaterials by the system nitric acid-potassium permanganate. 9 Functionalization of the surface resulted (Figure 6.1) in both a higher number of cells attached to the surface and changes in their morphology. Integrins, through which cells communicate with the surrounding environment, recognize the introduced changes and prove them by binding to the surface. As a result, it is possible to observe cells spreading and extending their filopodia in an oriented way after the oxidation. It should be noticed that there must be a compromise between functionalization and hydrophilicity. Proteins need some active places (in terms of charge and functionality) on the surface, where they can bind. On the other hand, the introduction of these active places is related to an increase in hydrophilicity. Generally, proteins have a hydrophobic nature and therefore repulsion but not adhesion can be observed when a surface with very high hydrophilicity is produced. (Actually, surface passivation with hydrophilic molecules is used for modification of devices in contact with blood. The passivated surface reduces or prevents the adhesion of thrombogenic cells and proteins onto the underlying substrate or material, thereby preventing surface-induced blood clotting.) After studying a wide variety of substrate polymers, Tamada and Ikada found21 that there is an optimal wettability for cell adhesion and that is approximately 70° water contact angle.

6.3.3

Hydrolysis

The ability of a material to be resorbed over time is an important property in many biomedical applications. Hydrolysis is the most common way through which the natural polymers degrade in the organism to normal metabolic compounds. All biomaterial surfaces are potentially susceptible to hydrolysis, simply due to the fact that they are surrounded by a warm aqueous environment (the body fluids) containing hydrolysing agents (e.g. enzymes). Catabolism of starch by α-amylase (Figure 6.2), which is available in the human blood © 2008, Woodhead Publishing Limited

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100 µm (b)

100 µm (c)

100 µm (d)

6.1 SaOs-2, cultured for seven days on SPCL (a and b) and SEVA-C (c and d) before (a and c) and after (b and d) surface oxidation by potassium permanganate. © 2008, Woodhead Publishing Limited

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100 µm (a)

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α-Amylase

OH

OH

O

O *

*

O OH

O OH

OH OH

OH

O

O

OH

OH OH

Amylose (starch)

OH O

O *

OH OH

OH O Reduced sugars

*

6.2 Enzymatic (α-amylase) hydrolysis of starch.

and in the saliva,22 is one example for those processes. Natural polymers containing ester, amide or other carboxylic derivative groups undergo degradation by a simple hydrolytic mechanism (Figure 6.3). The reaction is base- or acid-catalysed and sensitive to temperature above 37°C. Chitosan, a well known biomaterial for various applications, is produced from chitin (Figure 6.4) using this process. On the other hand, hydrolysis is a powerful surface modification method. More hydrophilic surfaces can be produced via the attack of a nucleophil agent.23–26 Sodium and potassium hydroxides are the most used nucleophils. The altered surface functionality can be used for further chemistry24–26 including immobilization of biomolecules.

6.4

Physical methods for surface alterations

6.4.1

Plasma activation and modification

Plasma is considered27 as the fourth state of matter (Figure 6.5). It contains various (atomic, molecular, ionic and radical) energetic, reactive, positively and negatively charged species but as a whole, plasma is neutral. The energy required to create and sustain plasma is supplied by an external electrical field. Various plasma sources can be used – gaseous (radio frequency glow discharge and corona discharge), metallic, and laser based. The plasma state exists only at a low pressure (less than 1–10–2 torr). Several plasma techniques are widely used for surface modification of natural based polymers:

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H O

HO

H

O

H

HO HO

6.3 Hydrolysis of esters catalyzed by acid (upper) or base (lower).

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H

O

O

O

O H

H

O

O

ROH

OR

O H

O

OR

OR

.. OR

O

H2O ..

O

O

OR

OR

H

H

O

O

OR

O

H

H

H

H

OH

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Chitin

OH

OH O O

O

O OH

*

* OH

NH

NH

O

O

CH3

CH3 KOH or NaOH, 100°C, 1–2 hrs

OH OH

O

O

O *

O OH

* OH

NH2

NH2 Chitosan

6.4 Hydrolytic process involved in the conversation of chitin into chitosan.

Ions and electrons move independently, large space

Kinetic energy

Molecules, free to move, large spacing

Dissociation

Molecules, free to move Molecules, fixed in lattice

Ionization

Vaporization

Melting

Melting point

Boiling point

6.5 Transitional states of matter.

• • •

Plasma sputtering and etching; Plasma functionalization; Plasma polymerization.

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All these plasma techniques have several advantages: (a) All processes are restricted to the topmost (ångström) layer and therefore the modified material has similar bulk chemical and physical properties to the original one; (b) Modification is fairly uniform over the whole surface even for samples with complex shapes; (c) Surfaces of all kinds of materials can be modified, regardless of their structures and chemical reactivity. How does it work? When the plasma comes into contact with the biomaterial surface, the activated species are accelerated towards the substrate by the applied electric field. Since some parts of the surfaces are exposed to energies higher than the bonding energy of polymers, these parts undergo chain scission. The chain scission process will initiate various chemical and physical events.2,28,29 Surface degradation can be observed with sufficient sputtering time and enough (different for different materials) high power applied. Figure 6.6. shows an example of how the conditions, used for the plasma treatment, can alter the surface morphology of a material. A blend of starch and cellulose acetate (50/50 %wt) was treated at different powers and for different times. As can be seen from the scanning electron microscopy (SEM)

(b)

(a)

5 µm

5 µm (d)

(c)

5 µm

5 µm

6.6 Effect of plasma working conditions on the surface morphology of starch/cellulose acetate (SCA) blend (50/50 wt%): SEM micrographs of untreated SCA (a); and Ar plasma modified SCA at 80W, 15 min (b); at 30W, 15 min (c) and at 80W for 5 min (d).

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micrographs, all modified samples presented much rougher surfaces compared to the original one. This effect depends on the power used, which determines the acceleration of the active species toward the material surface, as well as on the time during which the material is exposed to this bombardment with active species.30 Plasma etching can be used either for cleaning off the surface of the material or as a surface morphology modification technique. Engineering of new composites with improved adhesion between the components31 and surfaces with better biocompatibility30,32 are only two examples of the enormous benefits which surfaces with tailored roughness/ surface area can bring to the material sciences arena. On the other hand, chain scission results in the formation of highly reactive surface radicals. Those radicals can be used either in subsequent plasma depositions/polymerization processes 33,34 or they can recombine (e.g. crosslinking reactions) with the other active species available in the reactor. Additional to the power and the exposure time, the working atmosphere is of principal importance for these processes. Gases such as CH435 or CF430,36,37 are usually used to decrease the wettability of the surface. Contrary, the use of oxygen (introducing –OH, –C=O, –COOH groups) or nitrogen (–NO2, –NH2, –CONH2 groups) plasma is one of the most powerful methods for increasing material hydrophilicity which usually results in improved adhesion strength, biocompatibility, and other pertinent properties.29,38 Chitosan, modified by oxygen34 or nitrogen32 plasma displayed a higher number of cells attached to the surface (Figure 6.7) and a higher proliferation rate compared to the untreated chitosan membranes, for which next to no cell adhesion was observed. All these processes can be applied to three dimensional (3D) samples only if the holes/trenches are wider than the mean free path of the electrons and the Debye length.39 Only in this case will the discharge, which generates the active species, be sustained. Highly porous and interconnected starch based (starch/polycaprolactone 30/70 wt%) scaffolds were modified by oxygen plasma. Dramatic improvement of human umbilical vein cell (HUVEC) adhesion on the modified samples can be seen in Figure 6.8.

6.4.2

UV irradiation

UV irradiation resembles getting a tan under the sun, and the same rules are followed: time and intensity of the irradiation are important factors and ‘sunburn’ could be caused if they are not within limits. Similarly to plasma treatment, UV irradiation can result in chemical (photo-crosslinking, photooxidation in air, or photochemical reactions in reactive atmosphere) or physical (surface morphology, etc.) changes.38,40–42 These photochemical reactions can be surface-limited or can take place deep inside the bulk (unlike plasma) depending on the UV absorption coefficient at the specific UV-wavelength

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(a)

(b)

200 µm (c)

20 µm (d)

200 µm

20 µm

6.7 SEM micrographs showing the effect of oxygen plasma modification (30W, 15 min) on SaOs-2 adhesion (three days of culture): untreated chitosan membrane (a and b) and modified ones (c and d).

150 µm

150 µm

6.8 Immunostaining (PECAM, Phaloiedin and nuclei) of HUVEC cultured for seven days on SPCL untreated (left) fibre mesh and SPCL fibre mesh modified by oxygen plasma (right).

(Lambert-Beer’s law). There are two groups of sources: continuous wave (CW) UV-lamps with a moderate light, and pulsed laser. The laser sources cause mainly surface etching. They can be used to modify very small surface areas and this is the reason for their wide application in micro- and nano-

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fabrication technologies. The CW UV-lamps are used for surface oxidation.41,42 Starch-based biomaterials have been modified by CW UV-lamp. As expected, no significant effect on surface morphology was observed. The irradiation resulted in surface oxidation and a higher number of cells adhered to the surface (Figure 6.9).

6.4.3

β- and γ-irradiation

The perfect sterilization procedure for natural based biomaterials is one that does not include any changes in chemistry, mechanical properties or degradation behaviour. In other words, the final make-up of a biomaterial should not destroy all the work already done. Radiation with γ- or β-rays is often used to sterilize extracorporeal and intracorporeal medical devices made from polymers. High-energy radiation in addition to killing bacterial life, may also affect material properties. The surface is not an exception – surface chemistry and surface energy could be inadvertently altered by cleaning and sterilization procedures.43 Sometimes, the sterilization process can be used as a surface modification technique. For example, it was found44 that sterilization of membranes from chitosan-soybean protein isolate by βirradiation increases the surface energy but does not affect the bulk properties of the material. Unfortunately, not always does the synergy modification/ sterilization work out. Studies45,46 on the effect of gamma irradiation on collagen structure clearly indicate chain scission resulting in a fraction of lower molecular weight material. Material degradation leads to a loss of mechanical properties as well as to change in the surface roughness. Additionally, crosslinking could occur. Crosslinking reactions affect initial tensile strength (increase), surface hydrophilicity (decrease) and the properties related to these. In general, aromatic polymers are more resistant to highenergy radiation than aliphatic ones, while the presence of impurities and additives may enhance degradation and/or crosslinking.

6.5

Grafting

The main advantage of surface grafting is the long-term stability of the introduced chains onto the material surface. In contrast to physically coated polymer chains, in this method the chains are attached to the surface by covalent bonding which avoids their delamination.2,47 Many different synthetic routes can be employed to introduce graft chains onto the surface of polymeric substrates but generally, the grafting methods can be divided into two groups.48 Grafting-from methods utilize active species created on the polymer surfaces to initiate the polymerization of monomers (usually acrylic or vinyl) from the surface toward the bulk phase. In the case of grafting-to methods, the reactive species are carried by the preformed polymer chains, which are

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(b)

(a)

100 µm

(d)

(c)

100 µm

100 µm

6.9 Optical micrographs of osteoblast-like cells stained with methylene blue and cultured on untreated (a, c) and UV-irradiated (b, d) SCA (upper) and SPCL (lower) for seven days. © 2008, Woodhead Publishing Limited

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going to be covalently coupled to the surface. The fundamental step in grafting is the creation of reactive groups on the substrate surface. This could be done either chemically49–53 or more often by irradiation.54–56 The great majority of grafting processes involves a radical mechanism of polymerization of vinyl monomers. Plasma processes can be also used31,33–35,37,51 for surface functionalization via grafting. In this process, radicals created on the surface interact with monomers which can be introduced in the plasma reactor either as vapour or by pre-adsorption33 of the material surface. Alternatively, the surface can be pre-activated by plasma with subsequent immersion in the monomer solution. Some examples of the use of plasma treatment as a pre-activation technique are shown in Figures 6.10 and 6.11. A higher number of osteoblast-like cells, adhered to the surface of SPCL (starch/poly(ε-caprolactone), 30/70) after acrylic acid grafting, was observed. However, the cell did not show (Figure 6.10) the typical osteoblast morphology. When chitosan was modified in a similar fashion, cells were much more spread, with extended filopodias (Figure 6.11).

6.6

Bio-approaches: Mimicking the cell–cell interactions

As mentioned before, cells interact with a foreign device primary through proteins adsorbed onto the surface. Section 6.3 to 6.5 of this chapter described some methods for tailoring the protein adsorption and consequently the cell behaviour through modification/functionalization of the material surface. However, the described methods are quite general, i.e. they are not selective for a certain protein or cell type. On the other hand, the body fluids are rich

100 µm

100 µm

6.10 SPCL untreated (left) and surface modified by acrylic acid grafting (Ar plasma activation, right): Effect of the treatment on cell (SaOs-2) adhesion after seven days of culture – methylene blue staining.

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200 µm

20 µm

200 µm

20 µm

6.11 SEM micrographs of SaOs-2 cultured for seven days on untreated chitosan membranes (upper) and membranes grafted with vinyl sulfonic acid after oxygen plasma activation (lower).

in highly competitive protein molecules and very often, those which are not desired are ‘faster’ and cover the available space on the surface. How to overcome this problem and to engineer a selective surface? One of the approaches is to pre-immobilize an instructive component on the surface which will further direct cell behaviour. Carefully selected proteins, as a part of the communication system of the cell, can be used as interpreters which translate the desired surface-cell information. On the other hand, phospholipids are the main building part of different bio-membranes. Therefore, they can be useful in a strategy, aiming to dupe the cell. Several methodologies for mimicking these two cell components are described below.

6.6.1

Protein immobilization

Several different methodologies have been used in order to immobilize different proteins on the material surface. Coating with proteins can be achieved by a simple physical adsorption. Protein physical adsorption will occur when the change in Gibbs free energy of the system decreases during the adsorption process. Generally, proteins adhere to hydrophobic surfaces,57 because of their hydrophobic nature, and are repelled by hydrophilic surfaces. A comparative study between starch-based materials showed58 that the most hydrophilic blend (starch/cellulose acetate, 50/50 wt%, SCA) adsorbs less protein than the blend with the biggest water contact angle (starch-poly(ethylene vinyl alcohol, 50/50%, SEVA-C) in unitary (fibronectin or vitronectin) or

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complex proteins solution system. However, most natural available materials are rich in polar groups (–OH, –NH2, –COOH, –SO3H, etc.) and therefore relatively hydrophilic. How then can proteins be irreversible deposited on the natural materials’ surface? Fortunately, more of the natural polymers bear charges that can be used in physical protein adsorption. Chitosan is an example of a polycation and hyaluronic acid can illustrate a polyanion (Figure 6.12). Electrostatic interactions between charged peptide residues presented by a protein’s surface and surface functional groups greatly contribute to the Gibbs free energy of protein adsorption.59 The layer-by-layer technique (LbL)60–62 is based on these interactions and follows quite a simple procedure (Figure 6.13). Recently, it was reported63 that both the number of the deposited layers and the charge of the last layer influence the adsorption of fibronectin. Furthermore, modulation of HUVEC attachment on the natural polymers, modified by fibronectin adsorption by way of LbL technique, was achieved. When the surface does not bear a charge, pre-activation or pre-modification, using one of the already described techniques, and subsequent protein immobilization can be a solution. There are several examples when this steptreatment is very successful. Laminin was incorporated64 onto chitosan, preactivated by plasma or wet chemistry methods. Although a significant increase of cell attachment was observed for both cases, plasma treatment was indicated as a better methodology for the protein grafting on chitosan membranes. A similar effect was reported65 for starch-based biomaterials, pre-activated by plasma and subsequently immersed in different protein solutions. The use of whole proteins carries some disadvantages for application in the medical field. Proteins must be isolated from other organisms and purified. Otherwise they may elicit undesirable immune responses and increase infection risks. Normally they are expensive and often not available in a clinically acceptable form. Due their stochastic orientation on the surface, not all proteins have an appropriate orientation for cell adhesion.66 The incorporation of short oligopeptides having specific binding domains can overcome most of the indicated problems. The advantages of using small peptides rather than whole proteins are that they are relatively inexpensive to synthesize and easy to purify. Additionally, they exhibit higher stability to sterilization processes, heat treatment and pH variation, storage and conformation shifting and they can be characterized easily.67 Furthermore, when they are covalently bonded to the surface, they are more stable to cellular proteolysis than adsorbed cell adhesion proteins, since protein desorption is eliminated and the active groups are not exposed to soluble proteases. In 1984, Pierschbacher and Ruoslahti published a pioneer work,68 in which Arg-Gly-Asp (RGD) was identified as the first adhesive recognition sequence in fibronectin. Subsequently, the same motif was identified in other celladhesion proteins such as vitronectin, collagen or laminin.69 Nowadays, there are several short oligopeptides’ sequences used69–71 to mediate cell-specific

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O *

COO

OH

O

*

OH NH3

*

O O

OH

OH O

O HO

O

HO O

*

OH NH

NH3

O

C CH3

6.12 Two examples of natural polyions: chitosan (left) which is polycation at low pH and the polyanion hyaluronan (right).

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OH

Surface modification for natural-based biomedical polymers

Polyanion solution

Washing

Polycation solution

183

Washing

6.13 Schematic representation of layer by layer (LbL) deposition technique depicting film deposition starting with a positively charged substrate. Table 6.3 Some of the identified active sequences from various proteins and the receptors which recognized them Protein

Recognition sequence

Receptor

Fibronectin

Gly-Arg-Gly-Asp-Se (RGD) Leu-Asp-Val Arg-Glu-Asp-Va (REDV)

α5β1, αIIbβ3, ανβ3, α3β1, ανβ1 α 4β 1 α 4β 1

Laminin

Tyr-Ile-Gly-Ser-Arg (YIGSR) Pro-Asp-Ser-Gly-Arg (PDSGR) Arg-Tyr-Val-Val-Leu-Pro (RYVVLPR) Leu-Gly-Thr-Ile-Pro-Gly (LGTIPG) Arg-Gly-Asp (RGD) Ile-Lys-Val-Ala-Val (IKVAV)

67-kDa binding protein ? ? 67-kDa binding protein ? 110-kDa

Vitronectin

Arg-Gly-Asp (RGD)

ανβ3, ανβ5, αIIbβ3

Fibrinogen

Arg-Gly-Asp (RGD)

ανβ3, αIIbβ3

von Willebrand factor

Arg-Gly-Asp (RGD)

αIIbβ3

Entactin

Arg-Gly-Asp (RGD)

?

Collagen type I

Arg-Gly-Asp (RGD) Asp-Gly-Glu-Ala (DGEA)

30, 70, and 250 kDa α 2β 1

adhesion and function (Table 6.3). As in the grafting process, several different methodologies can be used in order to create a chemical bond between the oligopeptides and the surface of the material. Photografting of GRGD onto chitosan was reported72 to improve the adhesion and proliferation of endothelial cells on the modified surfaces. On the other hand, chemical methods can also be used. Carbodiimide chemistry is very often used73,74 as a strategy for protein conjugation. This strategy has several advantages. Either membranes or samples with complex geometry can be coated. Moreover, the reaction can be performed in aqueous or organic media by using different carbodiimides.

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Therefore, the solubility should not be an obstacle for the process. Taking advantage of the highly reactive chitosan amine group, GRGD was grafted74 on 3D chitosan structures. The peptide density on the surface was measured to be around 10–12 mol/cm2, promoting cell adhesion and proliferation as well as enhancing the formation of mineralized foci. Nevertheless, this reaction has a disadvantage. Two different acid moieties (end group on Ser and the carboxylic acid of Asp) in the RGD are present. This presence imposes additional step-protection of the acid group on Asp, without which control of the reaction is difficult. There is an alternative strategy, which uses succinic anhydride73 to generate carboxyl groups (but not amine) on the chitosan surface. The created carboxyl groups can further react with the free amine groups of the peptide forming the necessary spacing between the surface and the peptide.67 The same strategy has been applied75 for alginate hydrogels, which bring the carboxyl groups in their native structure and no additional transformation before the conjugation is needed. Besides the surface functionality, which determines the binding oligopeptidesurface, the surface concentration and distribution of the immobilized active sequence are other issues that need attention. The minimal RGD surface concentration necessary for maximal cell spreading is 1fmol/cm2.76 The formation of focal contacts and stress fiber was observed at 10fmol/cm2. These values were calculated for RGD peptide immobilized on a poorly adhesive glass substrate. On the other hand, Jin Li et al. confirmed73 a dependency on the concentration of the peptide immobilized on chitosan membrane surfaces. Higher peptide concentration enhances the process of cell attachment, proliferation, migration and mineralization. Finally, there are also some disadvantages of using short protein sequences. Loss of both affinity and specificity of the sequence when taken out of the context of the protein, are some of them. For example, the hexapeptide Gly-Arg-Gly-AspSer-Pro (GRGDSP), which is the active sequence from fibronectin, is 1000 times less effective.

6.6.2

Lipid coatings

Lipids are not always useless burden! On the contrary, in the biomedical field they are even covetable. There are several reasons for this: (a) The lipid bilayers are the major building blocks of biological membranes; (b) They have hemo-compatible and non-thrombogenetic properties; (c) The phospholipids can self-organize into specific supramolecular aggregates. A simple approach for generating membrane-mimetic surfaces is to create supported lipid mono- or bilayers at the surface of bulk materials.77 Various methodologies (Table 6.2) of self-assembling monolayers (SAMs), Langmuir-

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Blodgett technique or covalent binding can be applied. Similarly to the proteins, SAMs are used57,59,77,78 as models since they are well defined and organized structures. Langmuir-Blodgett technique (LB) is the main technique used for the formation of lipid mono- or multilayers on natural-based polymers.79–82 The principle of the LB is illustrated in Figure 6.14. Phospholipid bilayer formation on chitosan and agarose has been performed79 using LB. It was found that bilayer lipid membranes, cushioned by thin chitosan films, are more stable than agarose-cushioned membranes. Charge, which the chitosan poses, is most probably the reason for this stabilization. Molecular weight of the polymer used is another factor to be considered.83 Lipid coated vesicles but not membranes are also objects of great scientific interest77,84 because of their application as release systems. The cell membrane, which is built by phospholipids among other bioactive components, cannot ‘recognize’ the lipid vesicles and allow them to penetrate inside the cell and to deliver the target component which is previously loaded in the core of the vesicle. A phospholipid coating on plasmid DNA adsorbed starch-chitosan nanoparticles has been investigated84 in order to create a barrier between DNAse sensitive genetic material and body fluids. Such a system possesses both the surface properties of a liposome and the drug loading effectiveness of polymeric nanoparticles. Another example is the so-called synthetic biomimetic supra-molecular BiovectorTM (SMBVTM),85 which has been proven in preclinical and clinical evaluation to be a suitable candidate for the delivery

Spread lipid solution

Compressed lipid film

Lipid monolayer deposition

Lipid monolayer, formed on the polymer film

6.14 Illustration of the Langmuir-Blodgett technique.

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of nasal vaccines. In general, SMBVTM is a virus-like particle made of an inner core of polysaccharide hydrogel. It can be further surrounded by a lipid bilayer formed by ionic/hydrophobic interactions. Due to their bicompartmental structure, SMBVTM particles can be loaded with various active substances. All these studies show that fundamental biological processes can be successfully mimicked with the help of lipid coated natural materials.

6.7

Future trends

Using advances in the material sciences, biology and nanotechnology, we have learnt much from nature. These lessons imposed a shift towards third generation,5 resorbable nanostructured surfaces, enriched with specific biosignals, that once implanted will help the body heal itself. Nevertheless, we are still a long way from recreating the complexity and dynamics of the natural three-dimensional environment of cells, their ECM. It is likely that cells require the full context of this 3D nano-fibrous matrix to maintain their phenotypic shape and establish natural behaviour patterns. Achieving effective temporal control over the signals that are presented to cells in 3D artificial matrices is still a key challenge in optimization the outside-in signalling.

6.8

Acknowledgements

The authors acknowledge EU Marie Curie Actions, Alea Jacta EST for providing the PhD Grant to P. M. López-Pérez and the Portuguese Foundation for Science and Technology (FCT) for provide the postdoctoral grants to I. Pashkuleva (BPD/8491/2002). This work was also supported by The European Union funded STREP Project Hippocrates (NNM-3-CT-2003-505758) and the European NoE EXPERTISSUES (NMP3-CT-2004-500283).

6.9

References

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48 Bhattacharya A and Misra B N, ‘Grafting: a versatile means to modify polymers – Techniques, factors and applications’, Prog Polym Sci, 2004, 29(8), 767–814. 49 Amornchai W, Hoven V P and Tangpasuthadol V, ‘Surface modification of chitosan films-grafting ethylene glycol oligomer and its effect on protein adsorption’, Macromol Symp, 2004, 216, 99–107. 50 Shantha K L and Harding D R K, ‘Synthesis and characterization of chemically modified chitosan microspheres’, Carbohyd Polym, 2002, 48, 247–253. 51 Elvira C et al., ‘Plasma and chemical induced graft polymerization on the surface of starch based biomaterials aimed at improving cell adhesion and proliferation’, J Mat Sci: Mat Med, 2003, 14, 187–194. 52 Sashiwa H and Aiba S, ‘Chemically modified chitin and chitosan as biomaterials’, Prog Polym Sci, 2004, 29, 887–908. 53 Subramanian A, Rau A V and Kaligotla H, ‘Surface modification of chitosan for selective surface–protein interaction’, Carbohyd Polym, 2006, 66, 321–332. 54 Sawpan M A, Khan M A and Abedin M Z, ‘Surface modification of jute yarn by photografting of low-glass transition temperature monomers’, J Appl Polym Sci, 2003, 87, 993–1000. 55 Rajam S and Ho C C, ‘Graft coupling of PEO to mixed cellulose esters microfiltration membranes by UV irradiation’, J Membr Sci, 2006, 281(1–2), 211–218. 56 Mao C et al., ‘Various approaches to modify biomaterial surfaces for improving hemocompatibility’, Adv Coll Interf Sci, 2004 110(1–2) 5–17. 57 Prime K L and Whitesides G M, ‘Self-Assembled Organic Monolayers: Model Systems for Studying Adsorption of Proteins at Surfaces’, Science, 1991, 252, 1164–1167. 58 Alves C M, Reis R L and Hunt J A, ‘Preliminary study on human protein adsorption and leucocyte adhesion to starch-based biomaterials’, J Mat Sci: Mat Med, 2003, 14, 157–165. 59 Basalyga D M and Latour R A, ‘Theoretical analysis of adsorption thermodynamics for charged peptide residues on SAM surfaces of varying functionality’, J Biomed Mat Res, 2003, 64A(1), 120–130. 60 Decher G, ‘Fuzzy nanoassemblies: Toward layered polymeric multicomposites’, Science, 1997, 277, 1232–1237. 61 Tang Z et al., ‘Biomedical applications of Layer-by-Layer assembly: From biomimetics to tissue engineering’, Adv Mat, 2006, 18, 3203–3224. 62 Groth T and Lendlein A, ‘Layer-by-Layer deposition of polyelectrolytes – A versatile tool for the in vivo repair of blood vessels’, Ang Chem Int Ed, 2004, 43(8), 926– 928. 63 Wittmer C R et al., ‘Fibronectin terminated multilayer films: Protein adsorption and cell attachment studies’, Biomaterials, 2007, 28, 851–860. 64 Huang Y-C et al., ‘Surface modification and characterization of chitosan or PLGA membrane with laminin by chemical and oxygen plasma treatment for neural regeneration’, J Biomed Mat Res A, 2007, 82A, 842–851. 65 Alves C M et al., ‘Modulating bone cells response onto starch-based biomaterials by surface plasma treatment and protein adsorption’, Biomaterials, 2007, 28(2), 307–315. 66 Elbert D L and Hubbell J A, ‘Surface treatments of polymers for biocompatibility’, Ann Rev Mat Sci, 1996, 26, 365–394. 67 Hersel U, Dahmen C and Kessler H, ‘RGD modified polymers: biomaterials for stimulated cell adhesion and beyond’, Biomaterials, 2003, 24(24), 4385–4415.

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68 Pierschbacher M D and Ruoslahti E, ‘Cell attachment activity of fibronectin can be duplicated by small synthetic fragments of the molecule’, Nature, 1984, 309(5963), 30–33. 69 Yamada K M, ‘Adhesive Recognition Sequences’, J Biol Chem, 1991, 266(20), 12809–12812. 70 Shin H, Jo S and Mikos A G, ‘Biomimetic materials for tissue engineering’, Biomaterials, 2003, 24(24), 4353–4364. 71 Hubbell J A, ‘Biomaterials in tissue engineering’, Bio-Technology, 1995, 13(6), 565–576. 72 Chung T W et al., ‘Growth of human endothelial cells on photochemically grafted Gly-Arg-Gly-Asp (GRGD) chitosans’, Biomaterials, 2002 23(24), 4803–4809. 73 Li J et al., ‘Investigation of MC3T3-E1 cell behavior on the surface of GRGDScoupled chitosan’, Biomacromolecules, 2006, 7(4), 1112–1123. 74 Ho M H et al., ‘Preparation and characterization of RGD-immobilized chitosan scaffolds’, Biomaterials, 2005, 26(16), 3197–3206. 75 Rowley J A, Madlambayan G and Mooney D J, ‘Alginate hydrogels as synthetic extracellular matrix materials’, Biomaterials, 1999, 20(1), 45–53. 76 Massia S P and Hubbell J A, ‘An RGD spacing of 440 nm is sufficient for integrin alpha-V-beta-3-mediated fibroblast spreading and 140 nm for focal contact and stress fiber formation’, J Cell Biol, 1991, 114(5), 1089–1100. 77 Collier J H and Messersmith P B, ‘Phospholipid strategies in biomineralization and biomaterials research’, Ann Rev Mat Res, 2001, 31, 237–63. 78 Scotchford C A et al., ‘Protein adsorption and human osteoblast-like cell attachment and growth on alkylthiol on gold self-assembled monolayers’, J Biomed Mat Res, 2002, 59, 84–99. 79 Baumgart T and Offenhausser A, ‘Polysaccharide-supported planar bilayer lipid model membranes’, Langmuir, 2003, 19, 1730–1737. 80 Thompson M, Krull U J and Kallury K M, Lipid Membrane-based Device, US Patent 4824529, 1989, Allied-Signal Inc., USA. 81 Baumgart T and Offenhausser A, ‘Lateral diffusion in substrate-supported lipid monolayers as a function of ambient relative humidity’, Biophys J, 2002, 83, 1489–1500. 82 Ionov R et al., ‘Interactions of lipid monolayers with the natural biopolymer hyaluronic acid’, Biochimica et Biophysica Acta, 2004, 1667, 200–207. 83 Fang N et al., ‘Interactions of phospholipid bilayer with chitosan: effect of molecular weight and pH’, Biomacromolecules, 2001, 2, 1161–1168. 84 Baran E T and Reis R L, ‘Biomimetic approach to drug delivery and optimization of nanocarrier systems’, in Mozafari M R, Nanocarrier Technologies: Frontiers of Nanotherapy, Dordrecht, Springer, 2006, 75–86. 85 von Hoegen P, ‘Synthetic biomimetic supra molecular BiovectorTM (SMBVTM) particles for nasal vaccine delivery’, Adv Drug Deliver Rev, 2001, 51, 113–125. 86 Khan M A et al., ‘Effect of mercerization on surface modification of Henequen (Agave fourcroydes) fiber by photo-curing with 2-hydroxyethyl methacrylate (HEMA) ’, Polym-Plast Technol, 2005, 44(6), 1079–1093. 87 Pashkuleva I, Azevedo H S and Reis R L, ‘Surface structural investigation into starch-based biomaterials’, Macromol Biosci, 2008, 8, 210–219. 88 Vander Wielen L C et al., ‘Surface modification of cellulosic fibers using dielectricbarrier discharge’, Carbohyd Polym, 2006 65(2), 179–184.

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89 Zanini S et al., ‘Modifications of lignocellulosic fibers by Ar plasma treatments in comparison with biological treatments’, Surf Coat Tech, 2005, 200, 556–560. 90 Szymanowski H et al., ‘New biodegradable material based on RF plasma modified starch’, Surf Coat Tech, 2005, 200, 539–543. 91 Gubitz G and Cavaco-Paulo A, ‘New substrates for reliable enzymes: enzymatic modification’, Curr Opin Biotech, 2003, 14, 577–582. 92 Ogawa T et al., ‘Super-hydrophobic surfaces of layer-by-layer structured filmcoated electrospun nanofibrous membranes’, Nanotechnology, 2007, 18, 165–607. 93 Renneckar S et al., ‘Novel methods for interfacial modification of cellulose-reinforced composites’, ACS Symposium Series, 2006, 938, 78–96. 94 Pasquini D et al., ‘Surface esterification of cellulose fibers: Characterization by DRIFT and contact angle measurements’, J Coll Interf Sci, 2006, 295(1), 79–83. 95 Lindqvist J and Malmström E, ‘Surface modification of natural substrates by atom transfer radical polymerization’, J Appl Polym Sci, 2006, 100(5), 4155–4162. 96 Li J et al., ‘Investigation of MC3T3-E1 cell behaviour on the surface of GRGDScoupled chitosan’, Biomacromolecules, 2006, 7, 1112–1123. 97 Tsubokawa N and Takayama T, ‘Surface modification of chitosan powder by grafting of ‘dendrimer-like’ hyperbranched polymer onto the surface’, React Fun Polym, 2000, 43(3), 341–350. 98 Nurdin N, François N and Descouts P, ‘GRGDS-grafted chitosan for biomimetic coating’, in Domard A, Roberts G A F and Varum K M Advances in Chitin Science, Lyon, J. André Publishers, 1998, 378–383. 99 Roy D, ‘Controlled modification of cellulosic surfaces via the reversible addition – Fragmentation chain transfer (RAFT) graft polymerization process’, Austr J Chem, 2006, 59(3), 229. 100 Morao A et al., ‘Postsynthesis modification of a cellulose acetate ultrafiltration membrane for applications in water and wastewater treatment’, Env Progr, 2005, 24(4), 367–382. 101 Vilaseca F et al., ‘Chemical treatment for improving wettability of biofibres into thermoplastic matrices’, Compos Interface, 2005, 12(8–9), 725–738. 102 Zhang J et al., ‘Chemical modification of cellulose membranes with sulfo ammonium zwitterionic vinyl monomer to improve hemocompatibility’, Colloids Surface B, 2003, 30, 249–257. 103 Yuan J et al., ‘Improvement of blood compatibility on cellulose membrane surface by grafting betaines’, Colloid Surface B, 2003, 30(1–2), 147–155. 104 Thielemans W, Belgacem M N and Dufresne A, ‘Starch nanocrystals with large chain surface modifications’, Langmuir, 2006, 22(10), 4804–4810. 105 Zhou Q et al., ‘Xyloglucan and xyloglucan endo-transglycosylases (XET): Tools for ex vivo cellulose surface modification’, Biocat Biotrans, 2006, 24(1–2), 107– 120. 106 Kumar G, Smith P J and Payne G F, ‘Enzymatic grafting of a natural product onto chitosan to confer water solubility under basic conditions’, Biotech Bioeng, 1999, 63, 154–165. 107 Chao A et al., ‘Enzymatic grafting of carboxyl groups onto chitosan – to confer on chitosan the property of a cationic dye adsorbent’, Bioresource Technol, 2004, 91, 157–162. 108 Ogino A et al., ‘Surface amination of biopolymer using surface-wave excited ammonia plasma’, Jap J Appl Phys, 2006, 45(10B) 8494–8497.

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109 Abidi N and Hequet E, ‘Cotton Fabric Graft Copolymerization Using Microwave Plasma. II. Physical Properties’, J Appl Polym Sci, 2005, 98, 896–902. 110 Mukherjeea P, Jones K L and Abitoyec J O, ‘Surface modification of nanofiltration membranes by ion implantation’, J Membr Sci, 2005, 254, 303–310. 111 Kiatkamjornwong S, Mongkolsawat K and Sonsuk M, ‘Synthesis and property characterization of cassava starch grafted poly(acrylamide-co-(maleic acid)) superabsorbent via gamma-irradiation’, 2002, 43, 3915–3924. 112 Woo C K, Schiewe B and Wegner G, ‘Multilayered assembly of cellulose derivatives as primer for surface modification by polymerization’, Macromol Chem Phys, 2006, 207(2), 148–159. 113 Khan F and Ahmad S R, ‘Graft copolymerization and characterization of 2hydroxyethyl methacrylate onto jute fiber by photoirradiation’, J Appl Polym Sci, 2006, 101, 2898–2910. 114 Hassan M M, Islam M R and Khan M A, ‘Surface modification of cellulose by radiation pretreatments with organo-silicone monomer’, Polym-Plast Tech, 2005, 44(5), 833–846. 115 Yi H et al., ‘Biofabrication with Chitosan’, Biomacromolecules, 2005, 6(6), 2881– 2894. 116 Kongdee A, Bechtold T and Teufel L, ‘Modification of cellulose fiber with silk sericin’, J Appl Polym Sci, 2005, 96(4), 1421–1428. 117 Alves C M, Reis R L and Hunt J A, ‘Preliminary study on human protein adsorption and leukocyte adhesion to starch-based biomaterials’, J Mat Sci: Mat Med, 2003, 14(2), 157–165. 118 Hersel U, Dahmen C and Kessler H, ‘RGD modified polymers: biomaterials for stimulated cell adhesion and beyond’, Biomaterials, 2003, 24, 4385–4415. 119 Chung T-W et al., ‘Growth of human endothelial cells on different concentrations of Gly-Arg-Gly-Asp grafted chitosan surface’, Artificial Organs, 2003, 27(2), 155– 161. 120 Itoh S et al., ‘Effects of a laminin peptide (YIGSR) immobilized on crab-tendon chitosan tubes on nerve regeneration’, J Biomed Mat Res B, 2005 73(2), 375–382. 121 Taillac L et al., ‘Grafting of RGD peptides to cellulose to enhance human osteoprogenitor cells adhesion and proliferation’, Comp Sci Tech, 2004, 64(6), 827–837. 122 Morigaki K et al., ‘Photopolymerization of diacetylene lipid bilayers and its application to the construction of micropatterned biomimetic membranes’, Langmuir, 2002, 18, 4082–4089. 123 Fang N and Chan V, ‘Interaction of liposome with immobilized chitosan during main phase transition’, Biomacromolecules, 2003, 4, 581–588. 124 Yang F, Cui X and Yang X, ‘Interaction of low-molecular-weight chitosan with mimic membrane studied by electrochemical methods and surface plasmon resonance’, Biophys Chem, 2002, 99, 99–106. 125 Girod S et al., ‘Relationship between conformation of polysaccharides in the dilute regime and their interaction with a phospholipid bilayer’, Luminiscence, 2001, 16, 109–116. 126 Ye S H et al., ‘Design of functional hollow fiber membranes modified with phospholipid polymers for application in total hemopurification system’, Biomaterials, 2005, 26, 5032–5041. 127 Ye S H et al., ‘High functional hollow fiber membrane modified with phospholipid polymers for a liver assist bioreactor’, Biomaterials, 2006, 27(9), 1955–1962.

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7 New biomineralization strategies for the use of natural-based polymeric materials in bone-tissue engineering I. B. L E O N O R, S. G O M E S, P. C. B E S S A, J. F. M A N O, R. L. R E I S, 3B’s Research Group University of Minho, Portugal, and M. C A S A L, CBMA – Molecular and Environmental Biology Center, University of Minho, Portugal

7.1

Introduction

Materials scientists have much to learn from the way that nature assembles biologically important structures. Human bone results from a simple combination of inorganic and organic materials, but scientists have yet to produce a material capable of reproducing the structure of bone in its entirety. Nature is still the best material scientist when it comes to designing complex structures and controlling the intricate processing routes that lead to the final shape of living creatures. In designing new biomaterials for bone regeneration, surface properties must be modulated in order to mimic the tissue being replaced and then lead to the formation of new bone at the tissue/biomaterial interface. An ideal material for this type of application should possess mechanical properties matching those of the tissue being replaced, adequate degradation and biocompatible behaviour. Biodegradable polymers are potential biomaterials for this purpose, since they can be mechanically and biologically compatible with bone. Our group has proposed the use of naturally occurring polymers for tissue engineering since they can be tailored to retain their tissue supporting properties for given lengths of time and are gradually biologically degraded into non-toxic components that are absorbed by living tissues. This chapter describes the importance of using bone morphogenetic proteins in bone tissue engineering and the development of new calcium phosphate (Ca-P) coatings, used on substrates comprising polymers of natural origin, as a vehicle to deliver critical organic bone components that affect tissue response, such as growth factors to initiate osteoinduction.

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The structure, development and mineralization of bone

Bone, enamel and dentin are mineralized tissues found in vertebrates. Bone has a highly complex hierarchical structure with several levels of embedded structures that extend from the molecular, or micro, scale to the macroscopic scale (Mann, 2001; Rho, et al., 1998; Weiner, et al., 1999). At the microscale, collagen reinforced with apatite forms individual lamella that range in size from nm to µm, while at the macroscopic scale, interstitial bone is composed of osteons ranging in size from µm to mm (Rho, et al., 1998, Wang, 2003). This hierarchically organized structure has an irregular, yet optimised, arrangement and orientation of the components, making the bone material heterogeneous and anisotropic (Bonfield, et al., 1998).

7.2.1

Bone composition, structure and development

Bone, dentin, cementum and mineralized tendons, belong to a family of composite materials formed by mineral, collagen, water, noncollagen proteins, lipids, vascular vessels and cells. All the members of this family have mineralized collagen fibrils as the basic building block (Boskey, 1999a, Weiner, et al., 1998). The mineral present in this family of materials is calcium phosphate, in the form of apatite (Ca10(PO4)6(OH)2). Apatite crystals are associated with collagen I fibrils, from which the mineralized collagen fibrils originate. In bone, these crystals are extremely small, with an average length and width of about 500 × 250 Å and thickness of 20 to 30 Å, and they are probably the smallest crystals formed biologically (Lowenstma, et al., 1989, Weiner, et al., 1992). Their small size allows for easy incorporation and adsorption of new ions and also enables efficient and rapid dissolution from osteoclasts, which, together with osteoblasts, are the cells responsible for bone remodelling. Apatite crystals are plate-shaped, even though this mineral phase can have hexagonal crystal symmetry (Weiner, et al., 1986, Weiner, et al., 1998). Moreover, the crystals are intimately associated with collagen I, being aligned parallel to the long axis of the collagen fibrils in the organic matrix, which gives the bone its strength. Bone organic phase is made up of approximately 95% collagen type I and 5% non-collageneous proteins and proteoglycans (Marks, et al., 1996). The collagen structure is formed by two α1 polypeptide chains and one α2 polypeptide, 1000 amino acids long and about 80–100 nm in diameter, which form the collagen fibrils (Birk, et al., 1991; Boyde, 1972; Weiner, et al., 1992). The three polypeptide chains are wound together in a triple helix chain with a diameter of 1.5 nm and length 300 nm, where the amino acid chains are all parallel to each other and have their ends separated by a 35 nm space. Each fibril is separated from the neighbouring fibrils in such a way

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that a gap of 68 nm exists between the NH2-end of one triple-helical molecule and the COOH-end of the next triple-helical molecule (Hodge, et al., 1963, Weiner, et al., 1992). Transmission electron microscopy (TEM) showed that the plate-shaped HA crystals in the mineralized collagen fibril were arranged in traverse layers across the fibril (Traub, et al., 1989b, Weiner, et al., 1986). In the first stages of mineralization, the crystals grow inside the 68 nm gaps between the triple-helical fibril molecules and, as the crystals grow, they compress the collagen fibril molecules and eventually fuse together, forming plates of mineral. The way the mineralized collagen fibrils are organized into fibres results in different organizational patterns, including parallel, woven, plywoodlike and radial (Weiner, et al., 1998). The parallel fibril pattern is commonly observed in mineralized tendons and in parallel-fibred bone (Pritchard, 1956). This pattern is characterized by the fibrils being organized parallel to the long axis of the bone. Parallel distribution of the fibril arrays improves the mechanical performance of the bone in one specific direction (Weiner, et al., 1998). In the woven structure, the fibrils are joined together in bundles weakly packed and with poor orientation, and this structure is found in the skeletons of amphibians and reptiles (Weiner, et al., 1998). Woven bone is also common in mammalian embryos, later being replaced by other types of bone (Pritchard, 1956, Weiner, et al., 1998). The plywood-like structure is characterized by the presence of bundles, formed by fibrils parallel to each other, but oriented orthogonally in relation to the neighbouring bundles. This type of structure is characteristic of the cementum, which is a specialized bony substance covering the root of a tooth (Lieberman, 1993). This structure has isotropic properties which allows this type of bone to withstand compressive forces applied from different directions. Plywood-like structure is also found in lamellar bone, which is very common in mammals and is the most common type of bone in humans (Weiner, et al., 1998). A radial arrangement of collagen fibrils is characteristic of the bulk of dentin present in the inner layer of teeth. In this structure, the bundles of collagen fibrils are oriented randomly. The apatite crystals have two distinct arrangements. Those located inside the collagen fibrils have their c axis oriented parallel to the long axis of the fibril (Wang, et al., 1998). The crystals occupying the spaces between the collagen fibrils have a random orientation (Mishima, et al., 1986). It is suspected that these crystal arrangements give the dentin isotropic properties (Rasmussen, et al., 1976; Weiner, et al., 1998). Apatite and collagen I are two of the major components of bone, at around 5–10% of bone tissue (Boskey, 1999a; Weiner, et al., 1998). The third major component is water, which is located within collagen fibrils, and in the gaps between the triple-helical molecules (Weiner, et al., 1998). Water is important

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for the mechanical function of the mineralized collagen fibrils (Boskey, 1999b), besides being necessary for nutrition and the proper function of cells present in this tissue. These three major components are intimately associated into an ordered structure, the mineralized collagen fibril (Weiner, et al., 1998). Their proportions, however, can vary considerably between different family members of bone. Noncollagenous proteins make up about 5% of the dry weight of bone, and lipids approximately 2 to 8% of the organic matrix in bone tissue (Boskey, 1999b). The noncollagenous proteins include phosphoproteins and gammacarboxylated proteins. The phosphoproteins found in bone include osteopontin, bone sialoprotein, osteonectin and bone acidic glycoproteins (Boskey and Paschalis, 1999). These proteins are synthesized by osteoblast cells, and there are increased concentrations of these proteins at the mineralization front (Cowles, et al., 1998; Weinstock, et al., 1973). Osteocalcin is the major gamma-carboxylated protein present in bone tissue and is produced by osteoblasts during bone mineralization (Stein, et al., 1993). Also can be found different types of enzymes in bone, namely kinases, phosphatases and metalloproteinases, which are the three major families of enzymes that act at the bone matrix, and are also involved in mineral deposition. The kinases and phosphatases have antagonistic activities. Kinases are responsible for the phosphorilation of bone matrix proteins and phosphatases are involved in the dephosphorylation of these same proteins. Additionally, the phosphatase enzymes are responsible for the regulation of extracellular phosphate levels. The metalloproteinases enzyme family includes collagenases, gelatinases and proteoglycan-degrading enzymes, which are involved in the degradation of the extracellular matrix (Boskey, et al., 1999). Lipids are also present in bone tissue, as major components of cell membranes and as acidic phospholipids. The acidic phospholipids aggregate with calcium and phosphate ions and then complex with proteins, creating proteolipids. The levels of proteolipids increase just before calcification starts at the epiphyseal growth plate, a region where longitudinal bone growth occurs. High levels of these lipids are also detected in newly mineralized bone (Boskey, et al., 1980; Boskey, et al., 1996). Large proteoglycan molecules, such as aggrecan, epiphican and versican, are present in higher concentrations in non-mineralized regions compared with mineralized areas (Robey, et al., 1996). Small proteoglycan molecules are also components of the bone extracellular matrix (Fisher, et al., 1983), and are present in the form of chondroitin-sulfate proteoglycans such as decorin and biglycan (Boskey, et al., 1999). The bone cell structure is made up of a distinct population of cells, which are responsible for the maintenance of structural, biochemical and mechanical characteristics. This cell population is composed of four different cell types, namely osteoblasts, osteocytes, bone lining cells and

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osteoclasts (Lian, et al., 1999). Osteoblasts are responsible for the deposition of bone matrix and derive from mesenchymal cells, which in turn derive from the mesodermal cell layer in the embryo. Osteoclasts are multinucleated cells responsible for the resorption of bone tissue and result from the differentiation of cells of the hematopoietic system (Karsenty, 1999; Lian, et al., 1999; Olsen, et al., 2000). Both mesenchymal cells and hematopoietic cells are derived from stem cells present in bone marrow. The differentiation of both populations of cells into osteoprogenitor cells, from which osteoblasts originate, and into pre-osteoclasts, precursors of osteoclasts, is rigorously controlled during the process of bone formation and bone growth by a multistep event cascade involving a great number of different molecules (Lian, et al., 1999). Molecules such as cytokines, growth factors and hormones act as signalling molecules, inducing the proliferation of stromal mesenchymal cells, and thus giving origin to colonies of osteoprogenitor cells that grow and differentiate into pre-osteoblast cells (Lian, et al., 1999). Bone morphogenetic proteins (BMPs) are growth factors that, except for BMP-1, belong to the transforming growth factor β (TGF-β) family, a group of molecules that when applied locally can stimulate the formation of new bone tissue (Hogan, 1996). BMP2, BMP-4 and BMP-7 act as strong inductors of osteogenesis in vitro and in vivo (Asahina, et al., 1993; Wang, et al., 1990). The TGF-β group by itself is involved in osteoblast differentiation and in the synthesis of extracellular matrix by cells in vitro (Bonewald, 1996). More details regarding BMPs can be found in Section 7.3. Fibroblast growth factor (FGF) and insulin-like growth factor 1 (IGF-1) are also involved in the in vitro proliferation and differentiation of osteoprogenitor cells (Canalis, 1993, Canalis, et al., 1980, Rodan, et al., 1996). During the initial stages of osteogenic differentiation, the pre-osteoblast cells are responsible for the synthesis and organization of the bone extracellular matrix. These initial stages are characterized by the proliferation of preosteoblasts and by the expression of growth factors (TGF-β) and other proteins such as histone4, collagen, fibronectin and low levels of osteopontin. The proliferation stage is followed by a maturation period characterized by a modification in the composition of the bone extracellular matrix, where the proteoglycans versican and hyaluronan, produced by pre-osteoblasts, are replaced by the chondroitin sulphate proteoglycans decorin and biglycan, synthesized by osteoblasts (Robey, 1996). The expression of alkaline phosphatase reaches higher levels during the maturation period, and therefore proteins associated with mineral deposition (osteopontin, bone sialoprotein and osteocalcin) start to be secreted by osteoblasts. During these stages, the osteoblasts continue to produce collagen. The production and release of all these biomolecules is very important in the mineralization process (Lian, et al., 1999). As the mineralization process evolves, the osteoblasts become

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enclosed by the mineralized matrix and this induces morphological alterations in these cells, leading to differentiation into osteocytes (Pockwinse, et al., 1992). Osteocytes and osteoblasts maintain a permanent contact through a continuous exchange of information mediated by fluxes of calcium ions through gap junctions (Civitelli, et al., 1993, Donahue, et al., 1995; Lian, et al., 1999; Yamaguchi, et al., 1994). In mature bone, the osteocytes remain inside structures called canaliculi that resemble small capillary vessels. These structures result from the activity of the osteoclasts (Weiner, et al., 1998). The osteoclastogenesis process is controlled by hormones, cytokines and transcription factors. Pu.1 is the earliest transcription factor known to be involved in osteoclast differentiation. Other transcription factors that play critical roles in osteoclast differentiation are c-fos, NF-kB, c-src and mi (Karsenty, 1999, Lian, et al., 1999). Hormones, such as parathyroid hormone (PTH) and 1α,25-dihydroxyvitamin D3, are also important during osteoclast differentiation (Lian, et al., 1999). Osteoclasts, together with osteoblasts and osteocytes, are very important cells in the bone remodelling cycle. The remodelling process results from the coordinated activities of both osteoclasts and osteoblasts, and starts with the recruitment of mononucleated osteoclast precursors that fuse together, forming multinucleated pre-osteoclasts, and bind to the bone organic matrix, defining a circular, sealed zone. After the differentiation of osteoclasts is completed, these secrete protons and proteolytic enzymes into the circular zone, initiating a resumption process that is completed by mononucleated cells (Dempster, 1999; Murrills, et al., 1989). The tunnels resulting from osteoclast activity are refilled by osteoblasts, which deposit new bone (Weiner, et al., 1998). The formation of new bone is preceded by the deposition of a layer of cement, followed by layers of lamellar bone. At the end of the deposition process, a structure with an onion-like appearance is obtained, known as an osteon, which is formed by concentric lamellae of bone tissue. At the centre of the osteon is a blood vessel connected to the canaliculi system, where the osteocytes remain (Weiner, et al., 1998). Osteons are considered to be the basic structural unit of both cortical (compact) and cancellous (trabecular) bone. In compact bone, the osteon unit is called the Haversian system or cortical osteon, and in the trabecular bone it is referred to as packet or trabecular osteon (Dempster, 1999). Compact bone represents about 80–85% of the bone mass in adult humans. In this type of bone, the lamellae that incorporate the osteons are arranged in a concentric way, as described previously. In trabecular bone, the lamellae are organized in a more flattened way, sometimes following the curvature of the bone’s surface (Shea, et al., 2005). Trabecular bone is found in the epiphysis region of long bones and within flat bones. Anatomically, this type of bone is characterized by a network of trabecular that delineate a space filled with bone marrow (Carter, et al., 2001; Shea, et al., 2005).

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The classification of bones according to their architectural structure (trabecular vs. compact) can also be distinguished by the process of ossification during skeletal development (Boskey, 1999b). Intramembranous ossification occurs in the flat bone of the cranium and results from the direct differentiation of mesenchymal precursor cells into osteoblasts. Endochondral ossification occurs in most of the bones of the skeleton and, contrary to intramembranous ossification, is based on a cartilage template, cartilage anlagen (Olsen, et al., 2000). During endochondral ossification, the chondrocytes of the cartilage anlagen proliferate and generate hypertrophic chondrocytes, which start to secrete collagen X and vascular endothelial growth factors. These angiogenic growth factors induce the formation of blood vessels, which is known as angiogenesis (Erlebacher, et al., 1995; Olsen, et al., 2000). Angiogenesis, a physiological process, is followed by the migration of osteoblasts, osteoclasts and hematopoietic cells into the cartilage anlagen, inducing the establishment of ossification centres. As ossification proceeds, the osteoblasts gradually replace the cartilage matrix with bone matrix and the chondrocytes suffer apoptosis (Erlebacher, et al., 1995; Olsen, et al., 2000).

7.2.2

Mineralization of bone

In hard tissues, such as bone and mantle dentin, it is generally accepted that the mineralization process starts within extracellular-bond structures known as matrix vesicles (MV) (Anderson, 1967; Bonucci, 1967; Boskey, 1999b; Kirsch, et al., 1997b; Plate, et al., 1996; Wiesmann, et al., 2005). These MV are the result of a polarized budding process that occurs in osteoblasts, odontoblasts and in chondrocytes present in the epiphyseal growth plate, a region where longitudinal bone growth occurs. The MVs are about 50 to 200 nm in diameter, and their membranes are formed by a bilayer of phospholipids, similar to plasma membranes, which contain acidic phospholipids, phosphatidylserine and phosphatidic acid, which may act as captors of calcium ions during mineralization (Anderson, et al., 2005; Cotmore, et al., 1971; Peress, et al., 1974; Wuthier, 1975). A protein, annexin V, is found under the lipidic bilayer. Annexin V forms hexameric structures that surround hydrophilic pores through which calcium ions are transported into the MV (Kirsch, et al., 2000; Kirsch, et al., 1997a; Luecke, et al.; 1995, Nelsestuen, et al., 1999; Yang, et al., 2007). Annexin V is also responsible for the connection of MVs to collagen fibrils, where it functions as an anchor element (Kirsch, et al., 1992; Kirsch, et al., 1994). Another type of transporter also present in MV membranes is sodiumdependent phosphate, which is responsible for the inward movement of phosphorous into these vesicles (Anderson, et al., 2005; Montessuit, et al., 1995; Montssuit, et al., 1991). Matrix vesicles are rich in phosphatase enzymes,

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such as alkaline phosphatase (ALP), which is linked to the outer surface of the MV membrane (Harrison, et al., 1995, Matsuzawa, et al., 1971). Other enzymes present include adenosine monophosphoesterase (AMPase), adenosine triphosphoesterase (ATPase) and inorganic pyrophosphatase (PPiase). ALP and AMPase are responsible for the hydrolysis of adenosine monophosphate, a process that results in the release of inorganic phosphorous, which can be incorporated into calcium phosphate minerals as they grow. The enzymes PPiase and ATPase are involved in the hydrolysis of inorganic pyrophosphate and adenosine triphosphate, respectively, to release inorganic phosphorous (Ali, et al., 1970; Anderson, et al., 2005). Other non-collagenous bone-matrix macromolecular protein families can be found in the MV (Missana, et al., 1998), besides the enzymes mentioned above, including bone sialoprotein, osteonectin and osteocalcin, which might be involved in the control of nucleation and growth of the mineral phase (Boskey, 1998b; Boskey, et al., 2000; Kinne, et al., 1987). It is known that these proteins organize the extracellular matrix, control cell-cell and cellmatrix interactions, and provide signals to bone cells besides facilitating mineralization. As mentioned previously, it is generally accepted that mineralization starts inside the MV by binding calcium ions, which are transported by annexin V (Goldberg, et al., 1996; Wu, et al., 1997). The uptake of calcium ions is parallel to the inward movement of phosphorous ions, which occurs via the activation of sodium-dependent phosphate transporters. According to some studies, the phosphatidylserine first binds to calcium ions to form a calcium-phospholipid-phosphate complex (Anderson, 1967; Bonucci, 1967; Plate, et al., 1996). This binding with calcium ions, and phosphate ions, induces the formation of nucleation regions (Morris, et al., 1992; Plate, et al., 1996). When the critical radius for crystal nucleation is reached, a calcium phosphate mineral starts to precipitate as an amorphous structure. This amorphous structure is then converted into a hydroxyapatite (HA) mineral with a crystalline arrangement (Anderson, 1969; Anderson, et al., 2005; Gay, et al., 1978; Sauer, et al., 1988; Wu, et al., 1993). The MV becomes progressively filled with HA as it gets closer to the collagen fibrils (Mann, 2001; Sommerfeldt, et al., 2001). Then, at a point of supersaturation, mineral crystallization begins and, as the MV disintegrates, the mineralization nodule forms (Wiesmann, et al., 2005). At this point, the mineral is exposed to the matrix, where the subsequent crystallization takes place in association with collagen fibrils from the surrounding matrix (Calvert, 1994; Christoffersen, et al., 1991; Mann, 2001). Matrix vesicles may also locally remove pyrophosphate, an inhibitor of apatite crystal growth. In addition to local control of the levels of precipitating species and inhibitor, matrix proteins are also thought to act as nucleating sites (Calvert, et al., 1996). As a result, bone mineral forms within the collagen fibrils (Weiner ,

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et al., 1998), and this has been attributed to acid sites at the ends of the collagen triple helices (Calvert and Rieke 1996). The mineralization process first starts at the surface of the collagen fibrils and rapidly proceeds to the interior (Hohling, et al., 1980). This can be observed, at the initial stage of mineralization, in transversal sections of turkey tendon (Arsenault, 1988) and in embryonic fish dentin (Lees and Prostak, 1988), where crystals can only be seen at the surface of the collagen fibrils (Traub, et al., 1989a). These primary crystallites correspond to strands of apatite nodules, with a nanometer size, which are arranged parallel to the c-axis of collagen (Hohling, et al., 1980). As the mineralization stage progresses, the mineralization nodules continue to grow and eventually coalesce laterally with neighbour mineralization nodules and hence needle-shaped single crystals of HA are formed (Arnold, et al., 1997; Arnold, et al., 2001; Wiesmann, et al., 2005). These needle-shaped HA structures continue to grow and eventually fuse together, resulting in HA crystals with a ribbon or plate conformation (Hohling, et al., 1980). During this growing process, the highly ordered structure and parallel alignment with collagen fibres act as a template to control and determine the orientation of the HA crystals (Anderson, et al., 2005; Kirsch, et al., 1997c). Even though bone mineralization has been widely studied, some aspects remain unclear, in particular how the sequence of events is orchestrated so perfectly. Understanding the relationship between the matrix molecules and nucleation and mineral growth, requires a deeper knowledge of protein structure-function interactions.

7.3

Bone morphogenetic proteins in tissue engineering

In the human body, the formation of tissues is usually a well-orchestrated process with the interplay of cells and molecular messages mediated by proteins that have special functions, such as cytokines and growth factors. Due to donor scarcity, the risk of transplant rejection and the post-operational pain that occurs frequently in autografts (implants from the patient themselves), the tissue engineering approach is gaining momentum. Tissue engineering combines the use of scaffolds, stem cells and growth factors. Bone is one of the tissues with the most potential for self-regeneration and therefore is at the forefront of tissue engineering research. Bone morphogenetic proteins (BMPs) have sparked great interest in the field of regenerative medicine due to their specific and high potential for forming new bone (Reddi, 2005). They have been extensively researched over the last few decades in a quest to find late-stage tissue engineering products that might serve to regenerate bone in therapeutic applications.

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The discovery of BMPs – bone inductors

The history of BMPs dates back to 1889, when Senn noticed that decalcified bone could induce healing in bone defects (Senn, 1889). In fact, even earlier, in ancient Greece, Hippocrates questioned whether endogenous substances from the human body could work as therapeutic agents. Bone has long been recognized as one of the tissues in the human body that could regenerate. In the 1930s, Levander provided the first evidence of ectopic bone formation after injecting crude extracts of bone into muscle tissue (Levander, 1934, Levander, 1938). In 1965, Urist’s discovery that demineralized bone induced new bone formation when implanted in vivo marked a landmark in bone regeneration research. Urist named the protein component bone morphogenetic protein (Urist, 1965). Following this, Reddi attempted the isolation and identification of different BMPs. Reddi proposed that these agents were responsible for the differentiation of progenitor cells in the bone marrow to produce bone and cartilage cells, leading to bone regeneration (Reddi, 1981; Reddi, et al., 1972). It was not until 1988 that these proteins were individually identified and genetically reproduced (Wozney, et al., 1988). Thereafter, it was quickly discovered that the recombinant human bone morphogenetic protein 2 (rhBMP-2) could, by itself, induce the repair and regeneration of bone in different parts of the skeleton. In the years that followed, several preclinical trials have shown that BMPs efficiently stimulate bone growth along the spinal vertebrae, in craniofacial models and in long bone defects (Nakashima, et al., 2003; Seeherman, et al., 2005).

7.3.2

BMPs and biomineralization

The regeneration of bone is a remarkable and complex physiological process, and BMPs are among the most important biomolecules in this process (Reddi, 2005), making them potentially a useful clinical tool. BMPs play several important roles in developmental biology, during the formation of different tissues in the human body in a process called embryonic patterning (Kishigami and Mishina, 2005). Moreover, BMPs also play a role in the organogenesis of other tissues besides bone, such as BMP-2 in the heart (Callis, et al., 2005) and in neural tissues (White, et al., 2001), and BMP-7 in the kidney (Simic, et al., 2005) and in reproductive organs (Shimasaki, et al., 2004). The understanding of how molecular cascades of growth factors orchestrate cell differentiation and growth is of fundamental importance for designing novel tissue engineering products, for instance in the timed release of cocktails of BMPs and other morphogens from naturally occurring polymer matrices (Raiche, et al., 2004). Bone regeneration starts with an inflammatory phase during which various cytokines and growth factors are released into the injury site, attracting bone

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progenitor cells and prompting these to differentiate. Days after fracture, cells from the periosteum, the outer layer of connective tissue covering the surface of bone, replicate and form cartilage tissue and bone tissue, known as woven bone, both of which are later replaced by lamellar bone, restoring the original strength. This process is tightly regulated by BMPs, since both chondroblast and osteoblasts are required, and these must proliferate and differentiate in time- and space-specific ways during bone healing. Tissue engineering approaches therefore include the development of bi-layered scaffolds that attempt to bridge the defect, mimicking the natural process (Mano, et al., 2007), and possibly in future will deliver BMPs at the right times and in the right amounts. Different BMPs, such as those inducing bone (BMP-2, 4, 7) or those inducing cartilage (BMP-6, 12, 13, and 14), may be used in combination on polymeric scaffolds. BMPs are members of the TGF-β superfamily and bind to serine-threonine kinase receptors on the cell surface, triggering specific intracellular pathways that activate and influence gene transcription and have precise effects on cell proliferation and differentiation. The specificity of intracellular signals is mainly determined by type I receptors (Miyazono, et al., 2005). BMP binding acts through a pathway of SMAD signalling molecules, which are the main transducers of serine-threonine receptors and BMP signals. Different combinations of cell receptors provide different signals, which results in differences in cell phenotypes and tissue effects (Sebald, et al., 2004). Researchers are working to understand how using different BMP signals can regulate cell molecular biology and biochemistry. The complex involved consists of an activated receptor-regulated SMAD (R-SMAD) and a commonpartner SMAD (Co-SMAD) (Xu, et al., 2002), and triggers the activation or repression of several genes involved in cell differentiation or proliferation (Miyazono, 2000).

7.3.3

Recombinant BMPs for tissue engineering

Following the discovery of BMPs, purified BMPs were isolated from bone in an attempt to screen for their potential use in bone biomedical applications. Since BMPs were only extracted from bone in low amounts, researchers used recombinant technology to produce and purify these factors. Over the last two decades, the use of BMP produced by recombinant technology in tissue engineering products for bone has gained momentum. There are two types of recombinant expression systems, mammalian cells that allow us to obtain active protein but in low yields, and bacterial systems that produce much larger amounts of BMPs, but usually in insoluble inactive forms that require complicated refolding steps. The inherent difficulties in obtaining rapidly large amounts of bioactive BMPs makes them expensive, and so alternative approaches for producing them are necessary. One way of

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overcoming these limitations might lie in a new approach for producing large amounts of soluble and pure recombinant human BMPs being developed at the 3B’s Research Group. This method is based on a novel plasmid expression system in E. coli grown in a bioreactor, and demonstrates protein bioactivity in fat-derived human adult stem cells and in murine C2C12 cell lines (Bessa, et al., 2007). The research and development of novel avenues for obtaining recombinant BMPs (cloning, expression, purification and evaluation of their bioactivity and properties) will enable researchers to design much larger experiments involving growth factors and allow their use for pre-clinical and clinical tests.

7.3.4

The importance of BMPs in bone biomedical research

In Europe and the US, an estimated 5-10% of all bone fractures show deficient healing, leading to delayed union or non-union (Westerhuis, et al., 2005). These cause significant morbidity and stress to the patients and have financial implications. Advances in bone tissue engineering have led researchers to look for new strategies and devices to accelerate bone healing that could be used on a clinical basis. BMPs are, not surprisingly, of great interest, since these growth factors have been widely researched for clinical applications over the last two decades and recently received approval from the Food and Drug Administration (FDA) for human clinical use (Giannoudis, et al., 2005, McKay, et al., 2007). In 1997, recombinant human BMP-2 was used for the first time in spinal fusion patients. Eleven patients were treated with rhBMP-2 delivered via a collagen absorbable sponge, which was injected at the treatment site. Because the patients did not require bone grafting from the pelvis, their treatment was shorter and their post-surgical trauma was less than that typically seen in conventional bone grafting techniques (Boden, et al., 2000). Subsequently, BMPs have been studied extensively in several clinical trials and have received approval for human usage in cases of spinal fusion and long bone fractures (Boden, et al., 2002; Burkus, et al., 2003; Burkus, et al., 2006; Glassman, et al., 2007; Vaccaro, et al., 2003; Vaccaro, et al., 2005, Vaccaro, et al., 2004). Several areas of clinical application are currently under study, including spinal fusion and degenerative disc disease, long bone fractures and dental tissue engineering. There may be other clinical applications for BMPs, for instance craniomaxillofacial defects and diseases, improving osteointegration of metallic implants, musculoskeletal reconstructive surgery, tendon and ligament reconstruction, and periodontal and dental tissue engineering applications (Cheung, et al., 2006; Nakashima, et al., 2003). There is little doubt that powerful biological proteins such as rhBMP-2 will eventually help surgical specialists treat a variety of common bone defects and disorders.

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These osteoinductive factors will enable surgeons to modify their techniques to minimize the invasiveness of their operations. Ultimately, the goal is to reduce the pain associated with surgery and recovery, improve the effectiveness of surgical treatments, and hasten the return of patients to productive and healthy lifestyles. Spinal fusions comprise nearly half of all grafting surgery, and spinal fusion applications are an important part of current clinical trials (Carlisle, et al., 2005). The interest centres on the use of BMPs to accelerate healing in patients with degenerative disk disease, thus removing the need for autograft harvesting and reducing morbidity. Degenerative disc disease is defined as back pain caused by degeneration of the disc as confirmed by clinical data and symptoms. The approach is to use a collagen or other carrier soaked with BMP, which is implanted in the spine. Spinal fusion is performed in two different ways: posterolateral fusion, which involves placing the bone graft between the transverse processes in the back of spine, and interbody fusion, which involves placing the bone graft between the vertebrae in the area occupied by the intervertebral disc, usually by inserting the BMP in the spine through an anterior incision (from the front of the spine). There is little possibility for the growth factor to seep out and form bone where it is not needed, and clinical trials have shown this method to be very effective. Although most spinal fusion involves either collagen sponges or synthetic polymers, the use of naturally occurring scaffolds is currently being researched in animal models, including rats (Patel, et al., 2006). Fracture healing is another clinical situation where BMPs have been extensively researched for the development of possible tissue engineering products. Regeneration of bone fractures is a multi-stage cascade of events which involves the interaction of cells co-ordinated by complex signalling pathways. Treating long bone fractures using BMPs in combination with naturally occurring polymers is an active area of tissue engineering research. Tests on animal models include rabbit femurs, using fibrin hydroxyapatite composites (Sato, et al., 1991), bone defects in rats, using alginate gels (Saito, et al., 2005), and non-union tibial defects in rabbits, using hyaluronic acid gels (Eckardt, et al., 2005). In combination with collagen, there are reports of clinical trials in humans for treatment of long bone fractures with BMP-2 and BMP-7. Govender performed a trial with 450 patients with open tibial fractures. Patients were randomized to receive intramedullary nailing with different doses of rhBMP2 and after 12 months, results showed faster healing and reduced infection with higher doses of rhBMP-2 (Govender, et al., 2002). Friedlaender treated 122 patients with a total of 124 tibial non-unions in a randomized way, with patients receiving either the insertion of an intramedullary rod with BMP-7 in an absorbable collagen carrier or a bone autograft (Friedlaender, et al., 2001). The authors concluded that the method was a safe and effective alternative for tibial non-unions. In 2003, FDA approved the use of BMP-7

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in collagen sponges for treating long-bone non-unions as an alternative to autograft where this is unfeasible or contra-indicated. With the excitement over the potential clinical applications of BMPs, especially in novel delivery systems based on naturally occurring polymers, there is little doubt that in the near future BMPs will be part of regenerative medicine in bone and clinical traumatology. Given the evidence from animal studies, BMPs will probably play prominent roles in future tissue engineering products for treating patient fractures, non-unions and segmental defects, in spinal fusion and in periodontal approaches, possibly combined with the most recent advances in stem cell science, nanotechnology and genomics, and applied using biomimetic coated natural polymers (see Figure 7.1).

7.4

Bio-inspired calcium-phosphate mineralization from solution

New design and processing methods are needed for implants used in bone tissue engineering to promote fast tissue formation and integration within the body. An ideal material for use in bone replacement and regeneration applications should combine a mechanical performance matching that of the tissue to be replaced, adequate degradation and biocompatible behaviour. Scientists are using biomineralization strategies to develop new methodologies for designing novel functional materials. Designers of biomimetics systems for regenerating and replacing bone must remember that the physical structure of a biomaterial is a key factor in determining cellular response and hence dictates the range of biomedical applications for a particular material (Tan, et al., 2004). Biological responses such as the bone-bonding ability of the materials are very important for bone-related applications. It is essential that an implant shows bone-bonding behaviour, or osseointegration, through the formation of an apatite layer on the surface of the biomaterial (Kokubo, et al., 1990a). For example, calcium-phosphates (Ca-P) have been shown to be osteoconductive, where bone formation directed from the host bone towards the implant results in bonding (Yuan, et al., 2004). Osteoconduction highlights the possibility for guided bone formation on the biomaterial surface, and chemical bonds between newly formed bone and the biomaterial (Yuan, et al., 2004). Most of the available methods for producing adequate Ca-P coatings that are biocompatible and have osteoconductive surfaces capable of guiding bone formation (Clèries, et al., 2000; Kaciulis, et al., 1999; Leonor, 2003; Wei, et al., 1999; Yamashita, et al., 1994) have difficulty in controlling the Ca-P layer composition, degree of crystallinity and substrate bonding or adhesion ability (de Groot, 1998; Hayashi, et al., 1993; Tanahashi, et al., 1995). Plasma spraying (Aoki, 1991; Gross, et al., 1994) is the most commonly

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Bench Cloning Production of recombinant BMP

Expression Purification

Formulation (scaffold, particles) Development of a biomimetic coated natural polymer carrier for BMP

Bioactivity tests Incorporation/Release Bioactivity in vitro Bioactivity in vivo Tissue engineering construct

Isolation Stem cells research Culture conditions Differentiation

Human clinical trials/bone applications Bedside

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7.1 From bench to bedside: Strategy for a bone tissue engineering approach involving the use of recombinant BMPs, human stem cells and biomimetic coated natural origin polymers.

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used commercial technique for applying Ca-P coatings to implant surfaces and is approved by the Food and Drug Administration (FDA). This technique, however, has several disadvantages, including the formation of other phases, such as tricalcium phosphate (TCP) and calcium oxide, poor adhesion to the substrate, an inability to coat porous implants, and a restricted, line-of-sight application (de Groot, 1998; Shirkhanzadeh, 1991). In addition, the crystal structure of plasma-sprayed coatings is not uniform and the coatings consist of a mixture of crystalline and amorphous regions (Leeuwenburgh, et al., 2001). Despite the presence of multiple phases, the stable plasma sprayed HA coatings currently in use show little evidence of resorption up to nine months postoperatively (Vasudev, et al., 2004). If the Ca-P material is released from heterogeneous coatings, the resultant particles may initiate inflammation in surrounding tissues (Leeuwenburgh, et al., 2001). Plasma-sprayed coatings are thought to be susceptible to longterm failure at the implant interface and, as a consequence of coating failure, could produce HA debris, which, in turn, could result in osteoclast activation, bone loss, and aseptic loosening (Bloebaum, et al., 1994; Capello, et al., 1998; Dhert, et al., 1991). The coating therefore needs to exhibit long-term stability and at the same time act as a reservoir of calcium and phosphate ions for inducing increased bone formation and bonding (Fazan, et al., 2000). It has been claimed (Gledhill, et al., 2001) that to deliver better in vivo stability for long-term performance, the HA coatings should be highly crystalline, thus achieving a lower degradation rate as compared to amorphous or partly amorphous coatings. When the coating is applied in a biodegradable polymer, the combined materials should integrate within the tissues, and be progressively degraded and eventually fully replaced by bone material. The ideal implant should present a surface conductive to or that will induce osseointegration, regardless of implantation site or bone characteristics. (Puleo and Nanci, 1999).

7.4.1

Biomimetic calcium phosphate coatings

In science, the word biomimetics has been used by several authors from different perspectives (Abe, et al., 1990; Ball, 2001; Boskey, 1998a; Kokubo, et al., 1990a; Reis, et al., 1997a; Sarikaya, 1999; Sarikaya, et al., 2003; Stupp, et al., 1997). Based on this concept, Kokubo et al. (Abe, et al., 1990) developed a technique for coating different organic, inorganic and metallic materials with bioactive layers, and this has been designated as biomimetic coating. The main aim of this biomimetic process is to mimic biomineralization, leading to the formation of a bone-like carbonated apatite layer on the surface of a substrate. The methodology has been claimed to be very useful for producing highly bioactive and biocompatible composites with different mechanical properties (Kokubo, 1996; Kokubo, et al., 2001; Kokubo, et al., 2000; Kokubo, et al., 1999).

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The crystal size of a biomimetic coating is smaller and the structure more comparable to bone mineral than the large and sintered hydroxyapatite particles produced by plasma spraying (Leeuwenburgh, et al., 2001). Hence these bioactive layers have the capacity to develop interfacial mineralization much more rapidly than HA or TCP implants (Hench, 1988), and this bone-like apatite is supposed to provide a more favourable environment for bone cell seeding and proliferation than sintered HA (Yuan, et al., 2001). The original biomimetic coating methodology includes two steps which can be summarized as follows (Abe, et al., 1990, Hata, et al., 1995): the substrates are placed near CaO-SiO2-based glass particles (MgO 4.6, CaO 44.7, SiO2, 34.0 P2O5 16.2, CaF2 0.5 wt%) immersed in a simulated body fluid (SBF) (Kokubo, et al., 1990b) solution with ion concentrations nearly equal to those of human plasma (Na+ 142.0, K+ 5.0, Mg2+ 1.5, Ca2+ 2.5, Cl– 147.8, HCO 3– 4.2, 2– HCO 2– 4 1.0, SO 4 0.5 mM). The glass particles release large amounts of calcium and silicate ions, which are adsorbed onto the surface of the substrate to induce apatite nucleation. The calcium ions increase the degree of supersaturation with respect to apatite in the SBF, which accelerates apatite nucleation (this first period is described as the nucleation stage). To allow the growth of the apatite nuclei formed on the substrate in the first stage and the formation of an apatite layer, the substrate is immersed in another solution, e.g. 1.5 SBF with ion concentrations 1.5 times those of the SBF at 36.5°C (this second period is referred to as the growth stage). The thickness of the apatite layer increases as a function of immersion time, and the growth rate of the apatite layer increases with the increment of ion concentrations in the 1.5 SBF solution (Abe, et al., 1990; Kokubo, et al., 1990b). Although very popular and effective, the ‘traditional’ biomimetic process, using bioactive particles as nucleating agents, still presents some difficulties regarding the adhesion of the apatite layer to polymeric surfaces and for coating materials with complex shapes (Miyaji, et al., 1999). Reis et al. adapted the biomimetic methodology by rolling the samples on a bed of wet bioactive glass particles before immersion in the SBF solution (Reis, et al., 1997a). This method successfully coated different types of polymers and shapes, including a high-molecular polyethylene, a biodegradable starch, a poly (ethylene-co-vinyl alcohol) blend (SEVA-C) and polyurethane foam. However, problems associated with weak coating adhesion were still observed, although the results were better than for the original method. To overcome this problem, different surface treatments were tried on SEVA-C substrates (Oliveira, 2002; Oliveira, et al., 1999; Oliveira, et al., 2005) prior to immersion in SBF, including potassium hydroxide (KOH), acetic anhydride, UV radiation and overexposure to ethylene oxide sterilization (EtO). New biomimetic methodologies were then developed by the 3B’s Research Group based on different approaches, including impregnation with a sodium silicate gel (Oliveira, et al., 2002; Oliveira, et al., 2003b), pre-coating with

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a calcium silicate layer (Oliveira, et al., 2003a; Oliveira, et al., 2003b; Oliveira, et al., 2004) and incubation in supersaturated salt solutions (CaCl2, KCl and MgCl2) (Oliveira, 2002). These surface treatments were performed prior to immersion in an SBF, in order to generate nucleating sites for the formation of the apatite layers. The methodologies were aimed at: (a) reducing the incubation periods for apatite formation; (b) improving adhesion strength between the coating and substrate; (c) producing Ca-P layers with different (tailored) Ca-P ratios; and (d) coating the inside of pores in porous 3D architectures. This is also a simple and cost-effective way of producing CaP coatings and low processing temperatures mean that there is no adverse effect of heat on the substrates. This coating method has three important characteristics (Baskaran, et al., 1998): (a) the control of solution conditions, including ionic concentrations (supersaturation levels), pH, and temperature; (b) the use of functionalized interfaces to promote mineralization at the substrate surface; and (c) the formation of dense Ca-P films without the need for subsequent thermal treatments. An innovative coating methodology for producing an apatite layer has been proposed by Leonor and Reis et al. (Leonor, et al., 2003c), based on auto-catalytic deposition. This new approach uses a deposition route that does not require the use of electric current since it is based on redox reactions. Three types of solution are being studied, using alkaline and acid baths, to produce the novel auto-catalytic Ca-P coatings. This route seems to be a very promising and simple methodology for pre-implantation treatment to coat various types of materials prior to their clinical application. Recently, Tuzlakoglu et al. (Tuzlakoglu, et al., 2007) demonstrated that, using a simple biomimetic spraying methodology on chitosan fibre mesh scaffolds produced by wet-spinning, they were able to induce the formation of a Ca-P layer when immersed in an SBF. It is important to stress that, for all coatings, the final coating chemistry for an implant must be considered in relation to its application. In particular, the dissolution of Ca-P coatings, which plays an important part in the complex bone integration process, must be understood (Burke, et al., 2001). Ion release from Ca-P coatings may indirectly affect cellular processes involved in bone integration through altered ligand-cell receptor affinities, varied calcium and pH-dependent enzyme kinetics, and a compositionally or structurally altered extracellular matrix protein environment (Burke, et al., 2001, MacDonald, et al., 2001). The factors that affect ion release from thin-film coatings include Ca-P chemistry, coating roughness, and extent of coating strain (Burke, et al., 2001). The dissolution properties of Ca-P should therefore be adapted to the kinetics of osteogenesis (Daculsi, et al., 2002).

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Biomimetic coatings on natural-based polymeric substrates incorporating biomolecules as a carrier for delivering bone-related factors

The biomaterials currently used in the development of medical devices for implantation do not have any type of control over the biological response. Generally, host response to implanted biomaterials is stochastic and uncontrolled. This leads us to the conclusion that it is necessary to develop a new generation of biomaterials capable of controlling the host-implant interaction, which would result in a better biological response to the implanted device. Design strategies for creating biomimetic materials that direct the interaction with biological systems, such as the formation of tissue surrounding implants or regeneration within artificial matrices, have led to a new interdisciplinary field which can be described as molecular engineering (Healy, 1999; Healy, et al., 1999). The idea is for an organic substrate to act as a template, incorporating biologically active biomacromolecules that preferentially induce tissue formation consistent with the cell type seeded either on or within device. A large variety of biological functions could be built into the materials, such as incorporation of growth factors and cytokines to promote cell differentiation, enzymes to catalyse reactions and drugs for site-specific delivery. Biomimetic strategies are inspired by the natural mineralization process, where the minerals made by living organisms, usually composites of protein, polysaccharide and mineral, form under physiological conditions of temperature (37ºC) and pH (7.4). These conditions allow the incorporation of bioactive species without compromising their performance and improve the functionality of the inorganic layer at the implant interface. It is widely accepted that the biointegration of biomaterials involves a series of cellular and extracellular matrix events, some of which take place at the tissue-implant interface, and, in part, reflect the host response to the bulk and surface characteristics of the implanted material (Puleo, et al., 1999). Cells recognize synthetic materials via a complex protein over-layer that is formed on the material by adsorption from body fluids immediately after contact with the body (Daculsi, et al., 2002; Ratner, et al., 2004). This rather indirect relationship between material properties and cellular responses, mediated by the intervening protein layer, has complicated the development of biomaterials enormously (Daculsi, et al., 2002; Ratner and Bryant, 2004). Due to the complexities of the in vivo environment, the science of the bone/ implant interface is still not fully understood; in particular the role played by different biomolecules and their influence on initial bioadhesion, mineralization and coating dissolution, which has not received much attention (Bender, et al., 2000; Combes, et al., 2002). It is well known that protein adsorption constitutes one of the earliest

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events at the biomaterial-tissue interface, and this not only strongly influences the subsequent interactions of many different types of cells with the surfaces, but also determines the initial cellular response to the adsorbed surfaces (Horbett, et al., 1996; Lobel, et al., 1998). As proteins from biological fluids come into contact with synthetic surfaces, it has been hypothesized (Bender, et al., 2000; Lobel, et al., 1998) that cellular adhesion, differentiation and the production of extracellular matrix will be affected. It is known that proteins do more than facilitate mineralization. They organize the extracellular matrix, control cell-cell and cell-matrix interactions, and provide signals to the bone cells (Boskey, et al., 2000). Also, there are many additional enzymes, matrix proteins, and, of course, growth factors that contribute to the formation of bone and can induce specific cell and tissue responses, controlling the tissue-implant interface with molecules delivered directly to the interface (Boskey, et al., 2000; Puleo and Nanci, 1999). Studying their distribution, modification, and in vitro effects remains essential. Additionally, the manner in which the mineral is deposited, the orientation of the crystals and their size are influenced by proteins (Mei, et al., 1995). All these factors contribute to the strength of the mineralized tissue and stabilize the mineral content. In vivo, proteins play an important role in modifying and determining the physical and chemical properties of the tissue, and adsorbed proteins modulate cellular interactions that play an important role in hard tissue regeneration (Zeng, et al., 1999). Several studies investigating the influence of incorporated proteins and active enzymes on the formation of Ca-P coatings produced by biomimetic methods can be found in the literature (Areva, et al., 2002; Azevedo, et al., 2004; Azevedo, et al., 2005; Combes, et al., 2002; Combes, et al., 1999; Feng, et al., 2002; Leonor, et al., 2003a; Leonor, et al., 2005; Leonor, et al., 2004; Liu, et al., 2004; Liu, et al., 2005; Liu, et al., 2003b; Liu, et al., 2006; Liu, et al., 2001; Lu, et al., 2001; Luong, et al., 2006; Radin, et al., 1997; Vehof, et al., 2001; Wen, et al., 1999). This constitutes a novel approach to producing coatings with tailorable properties, which simultaneously exhibit controlled biomolecule release and bioactive behaviour, and is attractive because it can be used to control the release of biomolecules as a function of specific cell and tissue responses with time (Puleo, et al., 1999). In addition to osteoconductivity, Ca-P coatings have high affinity for proteins, which makes binding easier and also makes them ideal carriers for osteoinductive agents such as proteins (for instances, collagen, fibronectin, laminin, vitronectin), and osteogenic growth factors such as bone growth factors (BMPs), insulin-like growth factors (IGFs) and transforming growth factors (TGFs), which transform recruited precursor cells, thus initiating osteoinduction (Groeneveld, et al., 1999) and hence regeneration of hard tissues (LeGeros, 2002). As mentioned above, osteoconductive biomaterials are good materials for

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bone grafts, since they act as templates for bone formation and form a direct bond with bone. However, osteoconductive biomaterials only support bone regeneration passively; they are not able to stimulate bone formation (Liu, et al., 2004, Yuan, et al., 2004). Guided bone formation on osteoconductive biomaterial surface is limited in distance, and therefore osteoconductive biomaterials alone may not repair large bone defects. For large bone defect repair, bone formation far from the host bone bed should occur by osteoinduction (Yuan, et al., 2004). Osteoconduction is a kind of bone formation that does not start directly from osteogenic cells. It includes two steps, first cell differentiation from non-osteogenic cells to osteogenic cells, and second bone morphogenesis (Yuan, et al., 2004). When a biomaterial is implanted in a non-osseous site and induces bone formation, it is defined as an osteoinductive biomaterial (Yuan, et al., 2004). However, osteoinduction by Ca-P biomaterials is material dependent, i.e. there are several material factors which are relevant to osteoinductive potential, such as the three dimensional structure (Fujibayashi, et al., 2004; Yuan, et al., 2004). de Groot et al. (Liu, et al., 2003b; Liu, et al., 2001; Wen, et al., 1999) have demonstrated that bovine serum albumin can be successfully incorporated into the crystal lattice of mineral matrices coating metal implants when these are prepared by biomimetic co-precipitation of the relevant components. In addition, due to the degradation of these biomimetic coatings, protein molecules are released gradually (Liu, et al., 2001) rather than in a single rapid burst, as is the case with superficially adsorbed proteins, making such biomimetically prepared coatings of value as slow drug-release systems (Liu, et al., 2003b). Biomimetic co-precipitation is based on wet chemistry techniques, i.e. acid etching, incubation in boiling diluted alkali, precalcification, and immersion in a supersaturated calcification solution (Wen, et al., 1997, Wen, et al., 1998). This technique produces Ca-P coatings at physiological temperature (37°C), which has an important advantage over the conventional coating technique, plasma spraying, because osteogenic proteins can be coprecipitated in the coating and preserve their biological activity, creating a protein delivery system (Wen, et al., 1999). It has also been demonstrated that rhBMP-2 can be successfully coprecipitated with calcium phosphate on the surfaces of titanium alloy implants without loss of biological activity (de Bruijn, et al., 2000; Habibovic, et al., 2004; Liu, et al., 2004; Liu, et al., 2005; Liu, et al., 2003a; Liu, et al., 2006; Sun, et al., 2003). Similar work was reported by Vehof et al. (2001), where a titanium mesh coated with CaP and loaded with BMPs induced ectopic bone formation and also, due to the Ca-P coating, exhibited osteoinductive behaviour. Solution-phase growth enables the formation of calcium phosphate layers on implant surfaces, even porous surfaces. Bioactive proteins can be directly integrated in the structure of Ca-P coatings, maintaining a conformation

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close to their native form, which improves the functionality of the inorganic layer at the implant interface. Therefore, it can be said that the simplest biomimetic approach involves the design of single component systems that mimic the chemistry of the targeted biological material. The 3B’s Research Group (Azevedo, et al., 2004; Azevedo, et al., 2005; Leonor, et al., 2003a; Leonor, et al., 2005; Leonor, et al., 2004) has also been investigating the development of Ca-P coatings produced by biomimetic routes in SBF, where proteins are co-precipitated with the inorganic elements. The main aims of this approach are to produce Ca-P coatings on natural origin polymers with novel properties (in terms of morphology, crystallinity, stability, mechanical strength) and create a delivery system for therapeutic agents. In the last 12 years, starch-based polymers have been proposed by our group (Boesel, et al., 2004a; Boesel, et al., 2004b; Elvira, et al., 2002; Gomes, et al., 2002; Gomes, et al., 2006; Gomes, et al., 2003; Malafaya, et al., 2006; Mano, et al., 2003; Mano, et al., 2004; Marques, et al., 2002; Mendes, et al., 2003; Mendes, et al., 2001; Reis, et al., 1995; Reis, et al., 2000; Reis, et al., 1996; Reis, et al., 1997b; Salgado, et al., 2004; Salgado, et al., 2005; Sousa, et al., 2000; Sousa, et al., 2002; Vaz, et al., 2001) as alternative biomaterials for temporary biomedical applications. One of the main advantages of these materials for bone-related applications is the combination of mechanical performance with degradation behaviour (Azevedo, et al., 2003; Mano, et al., 2000; Mano, et al., 2004; Reis, et al., 1996; Reis, et al., 1997b; Sousa, et al., 2002; Vaz, et al., 2001). Additionally, it has been shown (Gomes, et al., 2001; Marques, et al., 2002; Marques, et al., 2003; Marques, et al., 2005a; Marques, et al., 2005b; Mendes, et al., 2003; Mendes, et al., 2001; Reis, et al., 2000; Reis, et al., 1996; Salgado, et al., 2004; Salgado, et al., 2005) that these materials can comply with the biocompatibility requirements of a biomaterial, as defined in international standards, which is not typical of biodegradable systems. Compared to other biodegradable polymers on the market, starch-based blends are the cheapest, and are available in much larger quantities from several renewable plant sources. A major advantage of starch-based polymers is the possibility of controlling their surface properties to facilitate the interaction between the modified material and the biological system (Demirgoz, et al., 2000). Nevertheless, in terms of bone bonding, these starch-based polymers cannot induce the formation of an apatite layer without a bioactive coating or the presence of bioactive fillers (Leonor, et al., 2003b; Leonor, et al., 2002a; Leonor and Reis, 2003; Leonor, et al., 2002b; Leonor, et al., 2004; Oliveira, et al., 1999, Oliveira, et al., 2003b; Oliveira, et al., 2005; Pashkuleva, et al., 2005). Another approach that has been studied in our group is the incorporation of specific hydrolytic enzymes with the Ca-P coatings that degrade the substrate, which presents an alternative strategy for controlling the degradation rate of polymeric biomaterials. A self-regulated degrading system would be particularly

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useful in tissue engineering scaffolding, allowing the growth of new tissue within the degrading construct. To our knowledge, tailoring the degradation kinetics of biodegradable biomaterials is a completely novel approach. The proteins used in these studies were: (a) bovine serum albumin (BSA), which was used as a model protein in order to simulate more closely the conditions found in vivo, since albumin is one of the first proteins to interact with an implanted foreign body; and (b) α-amylase, a starch-degrading enzyme, which was used to tailor the degradation rate of the starch-based biomaterial. The results of our work showed that protein molecules can be efficiently incorporated into biomimetic Ca-P coatings and preserve their enzymatic activities, as demonstrated by the release of reducing sugars from starchbased polymers coated with Ca-P films incorporating α-amylase enzyme (Azevedo, et al., 2005, Leonor, et al., 2005). Using the biomimetic route, we were able to apply the methodology to several materials, both synthetic and natural polymers, and of different shapes and sizes. A delivery system for BMPs as part of tissue engineering constructs for bone biomedical applications has been researched. The main role of a delivery system for BMPs is to keep these growth factors at the site of injury for a prolonged period of time, providing an initial support for the attachment of cells (Li, et al., 2001; Seeherman, et al., 2005). Our research group has investigated the incorporation of rhBMP-2 onto Ca-P coatings on natural polymeric substrates, namely 3D architectures, carried out in SBF and produced by biomimetic routes, as described above. Figure 7.2 shows micrographs, obtained by scanning electron microscopy (SEM), of Ca-P coatings grown on the surface of a starch-based polymer (30/70 wt% polymeric blend of corn starch with polycaprolactone, designated as SPCL) under different conditions. It can be seen that, after seven days immersion in 1.5x SBF during the growth stage, the surface of SPCL was covered with a dense and uniform Ca-P film. It is very important that the distribution of the Ca-P coating along the fibres does not compromise the overall morphology and interconnectivity of the 3D-fibre mesh scaffolds. At higher magnifications, a finer structure where needle-like crystals are agglomerated can be seen. Figure 7.2 shows that, using this methodology, we are able to coat threedimensional structures with a dense Ca-P film at a thickness around 5 µm without compromising the interconnectivity of the scaffold. However, due to the complexity of the system and to the number of variables involved, further studies need to be carried out. Our work demonstrated that it is possible to incorporate bioactive proteins such as rhBMP-2 through a biomimetic calciumphosphate coating technique. It therefore opens new possibilities for incorporating other bioactive agents, such as growth factors or specific enzymes, in order to induce a cellular response or other desired effect. Our group is currently conducting several studies to explore this further.

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(a)

(c)

250 µm

250 µm

(d)

(b)

25 µm

25 µm

7.2 SEM micrographs of the Ca-P coatings on the surfaces of SPCL after 7 days in SBF (growth stage): (a) SPCL control (without rh-BMP2); (c) 50 mg/mL rh-BMP-2, added in the growth stage. Magnification (b,d) showing a detail of the structure presented in (a,c). Crosssection of the Ca-P coating on the control (e) showing a thickness around 5 mm and a finer structure where needle-like crystals are agglomerated.

7.5

General remarks and future trends

Scientists are committed to finding materials suitable for regenerating tissues such as skin, cartilage, bone, blood vessels, nerve and liver using polymeric devices. The greatest promise for achieving dramatic improvements in longterm clinical repair of the skeletal system is to concentrate research efforts on creating a new generation of biomaterials that enhance the human body’s own repair mechanisms. New biomineralization strategies using biomimetic approaches could be a significant breakthrough in the bone replacement and regeneration field. Biodegradable materials, including those based on naturally occurring polymers, coated with biomimetic Ca-P layers and incorporating growth factors, may constitute an effective way to provide osteoconductive and osteoinductive properties in a single material. As the Ca-P layer undergoes degradation in vivo, the proteins will be released gradually, enhancing the potential of these coatings to serve as a slow-release carrier system for the delivery of growth factors at the implantation site. © 2008, Woodhead Publishing Limited

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The main challenge for the future is to engineer a new hybrid material with a controlled in vivo dissolution rate, essential for providing more effective carriers for osteogenic factors and expediting the osseointegration of the implant. Ideally, the carrier should be resorbed at a rate equal to that of bone formation. In fact, the authors believe that surfaces with biomimetic coatings, into which osteogenic growth factors are incorporated, hold great potential for use in clinical orthopaedics and dentistry to improve the regeneration of bone tissue, and thus expedite the reestablishment of full functionality at the implantation site. The next generation of biomaterials will also include materials that are designed to mimic existing biological materials, including self-assembled biomaterials, capable of self-organization into structures with different hierarchical levels, and biomimetic biomaterials, produced through the combination of calcium phosphates with synthetic or natural polymers.

7.6

Acknowledgments

I. B. Leonor thanks the Portuguese Foundation for Science and Technology (FCT) for providing her a PhD scholarship (SFRH/BD/9031/2002) and the European Union funded STREP Project HIPPOCRATES (NMP3-CT-2003505758) and the European NoE EXPERTISSUES (NMP3-CT-2004-500283).

7.7

References

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T. (1995), ‘Apatite coated on organic polymers by biomimetic process: improvement in its adhesion to substrate by glow-discharge treatment’, J Biomed Mater Res, 29(3), 349–57. Traub W., Arad T. and Weiner S. (1989a), ‘Three-dimensional ordered distribution of crystals in turkey tendon collagen fibres’, Proceedings of the National Academy of Sciences, 86, 9822–9826. Traub W., Arad T. and Weiner S. (1989b), ‘Three-dimensional ordered distribution of crystals in turkey tendon collagen fibres’, Proc Natl Acad Sci USA, 86(24), 9822– 9826. Tuzlakoglu K. and Reis R.L. (2007), ‘Formation of bone-like apatite layer on chitosan fibre mesh scaffolds by a biomimetic spraying process’, Journal of Materials Science: Materials in Medicine, 18(7), 1279–1286. Urist M.R. (1965), ‘Bone: formation by autoinduction’, Science, 150(698), 893–899. Vaccaro A.R. et al. (2003), ‘A pilot safety and efficacy study of OP–1 putty (rhBMP-7) as an adjunct to iliac crest autograft in posterolateral lumbar fusions’, Eur Spine J, 12(5), 495–500. Vaccaro A.R. et al. (2005), ‘A 2-year follow-up pilot study evaluating the safety and efficacy of op-1 putty (rhbmp-7) as an adjunct to iliac crest autograft in posterolateral lumbar fusions’, Eur Spine J, 14(7), 623–629. Vaccaro A.R. et al. (2004), ‘A pilot study evaluating the safety and efficacy of OP–1 Putty (rhBMP-7) as a replacement for iliac crest autograft in posterolateral lumbar arthrodesis for degenerative spondylolisthesis’, Spine, 29(17), 1885–1892. Vasudev D.V., Ricci J.L., Sabatino C., Li P. and Parsons J.R. (2004), ‘In vivo evaluation of a biomimetic apatite coating grown on titanium surfaces’, J Biomed Mater Res A, 69(4), 629–636. Vaz C.M., Reis R.L. and Cunha A.M. (2001), ‘Degradation model of starch-EVOH plus HA composites’, Materials Research Innovations, 4(5–6), 375–380. Vehof J.W.M., Mahmood J., Takita H., van’t Hof M.A., Kuboki Y., Spauwen P.H.M. and Jansen J.A. (2001), ‘Ectopic bone formation in titanium mesh loaded with bone morphogenetic protein and coated with calcium phosphate’, Plastic and Reconstructive Surgery, 108(2), 434–443. Wang M. (2003), ‘Developing bioactive composite materials for tissue replacement’, Biomaterials, 24(13), 2133–2151. Wang E. et al. (1990), ‘Recombinant human bone morphogenetic protein induces bone formation’, Proceedings of the National Academy of Sciences, 87, 2220–2224. Wang R. and Weiner S. (1998), ‘Human root dentin: Structural anisotropy and Vickers microhardness isotropy’, Connective Tissue Research 39(4), 269–279. Wei M., Ruys A.J., Swain M.V., Kim S.H., Milthorpe B.K. and Sorrell C.C. (1999), ‘Interfacial bond strength of electrophoretically deposited hydroxyapatite coatings on metals’, Journal of Materials Science: Materials in Medicine, 10(7), 401–409. Weiner S. and Traub W. (1986), ‘Organization of hydroxyapatite crystals within collagen fibrils’, Federation of European Biochemical Societies Letters, 206(2), 262–266. Weiner S. and Traub W. (1992), ‘Bone structure: from ångstroms to microns’, Journal of the Federation of American Societies for Experimental Biology, 6(3), 879–885. Weiner S., Traub W. and Wagner H.D. (1999), ‘Lamellar bone: structure-function relations’, J Struct Biol, 126(3), 241–55. Weiner S. and Wagner H.D. (1998), ‘The material bone: Structure mechanical function relations’, Annual Review of Materials Science, 28, 271–298. Weinstock M. and Leblond C.P. (1973), ‘Radioautographic visualization of the deposition

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of a phosphoprotein at the mineralization front in the dentin of the rat incisor’, Journal of Cell Biology 56(3), 838–845. Wen H.F., de Wijn J.R., Cui F.Z. and de Groot K. (1998), ‘Preparation of calcium phosphate coatings on titanium implant materials by simple chemistry’, Journal of Biomedical Materials Research, 41(2), 227–236. Wen H.B., de Wijn J.R., van Blitterswijk C.A. and de Groot K. (1999), ‘Incorporation of bovine serum albumin in calcium phosphate coating on titanium’, Journal of Biomedical Materials Research, 46(2), 245–252. Wen H.B., Wolke J.G., de Wijn J.R., Liu Q., Cui F.Z. and de Groot K. (1997), ‘Fast precipitation of calcium phosphate layers on titanium induced by simple chemical treatments’, Biomaterials, 18(22), 1471–1478. Westerhuis R.J., van Bezooijen R.L. and Kloen P. (2005), ‘Use of bone morphogenetic proteins in traumatology’, Injury, 36(12), 1405–1412. White P.M., Morrison S.J., Orimoto K., Kubu C.J., Verdi J.M. and Anderson D.J. (2001), ‘Neural crest stem cells undergo cell-intrinsic developmental changes in sensitivity to instructive differentiation signals’, Neuron, 29(1), 57–71. Wiesmann H.P., Meyer U., Plate U. and Höhling H.J. (2005), ‘Aspects of collagen mineralization in hard tissue formation’, International Review of Cytology, 242, 121– 156. Wozney J.M., Rosen V., Celeste A.J., Mitsock L.M., Whitters M.J., Kriz R.W., Hewick R.M. and Wang E.A. (1988), ‘Novel regulators of bone formation: molecular clones and activities’, Science, 242(4885), 1528–1534. Wu L., Genge B.R., Dunkelberg D.G., LeGeros R.Z., Concannon B. and Withier R.E. (1997), ‘Physicochemical characterization of the nucleation core matrix vesicles’, Journal of Biological Chemistry, 272, 4401–4411. Wu L.N., Yoshimori T., Genge B.R., Sauer G.R., Kirsch T., Ishikawa Y. and Wuthier R.E. (1993), ‘Characterization of the nucleational core complex responsible for mineral induction by growth plate cartilage matrix vesicles’, Journal Biological Chemistry, 268(33), 25084–25094. Wuthier R.E. (1975), ‘Lipid composition of isolated epiphyseal cartilage cells, membranes and matrix vesicles’, Biochimica et Biophysica Acta 409(1), 128–143. Xu R.H., Chen X., Li D.S., Li R., Addicks G.C., Glennon C., Zwaka T.P. and Thomson J.A. (2002), ‘BMP4 initiates human embryonic stem cell differentiation to trophoblast’, Nat Biotechnol, 20(12), 1261–1264. Yamaguchi D., Ma D., Lee A., Huang J. and Gruber H. (1994), ‘Isolation and characterization of gap junctions in the osteoblastic MC3T3–E1 cell line’, Journal of Bone and Mineral Research 9(6), 791–803. Yamashita K., Arashi T., Kitagaki K., Yamada S., Umegaki T. and Ogawa K. (1994), ‘Preparation of apatite thin films through rf-sputtering from calcium phosphate glasses’, Journal of the American Ceramic Society, 77, 2401–2407. Yang L., Zhang Y. and Cui F.Z. (2007), ‘Two types of mineral-related matrix vesicles in the bone mineralization of zebrafish’, Biomedical Materials, 2(1), 21–25. Yuan H. and de Groot K. (2004), ‘Calcium phosphate biomaterials: an overview’, in Reis and Weiner, Learning from Nature How to Design New Implantable Biomaterials: from Biomineralization Fundamentals to Biomimetic Materials and Processing Routes, Dordrecht, Kluwer Academic Publishers, 37–57. Yuan X., Mak A.F. and Li J. (2001), ‘Formation of bone-like apatite on poly(L-lactic acid) fibres by a biomimetic process’, J Biomed Mater Res, 57(1), 140–150. Zeng H., Chittur K.K. and Lacefield W.R. (1999), ‘Analysis of bovine serum albumin adsorption on calcium phosphate and titanium surfaces’, Biomaterials, 20(4), 377–384. © 2008, Woodhead Publishing Limited

8 Natural-based multilayer films for biomedical applications C. P I C A R T, Université Montpellier, France

8.1

Introduction

In the field of biomaterials, controlling the surface properties of the materials may be a means to influence cell behavior including recolonization, adhesion, migration or even differentiation. Therefore, various strategies have been developed to modify the materials surface properties, such as LangmuirBlodgett deposition and self-assembled monolayers.1 For about ten years, polyelectrolyte multilayer (PEM) coatings have emerged and become a new and general way to modify and functionalize surfaces whose applications range from optical devices to biomaterial coatings.2,3 The technique is based on the alternate deposition of polyanions and polycations.4,5 In recent years, the use of natural polyelectrolytes and biopolymers has emerged.6–9 On account of their biocompatibility and non-toxicity, these latter films constitute a rapidly expanding field with great potential applications: preparation of bioactive and biomimetic coatings,7,9,10 preparation of drug release vehicles,8,11 buildup of cell adhesive or anti-adhesive films,6,9 and more recently creating a membrane mimetic barrier for islet encapsulation.12 In addition, natural polymers are already widely used for biomedical applications including hydrogel preparation, soft tissue repair, 13,14 drug delivery, 15 and viscosupplementation.16 Among the polysaccharides that are often used for biomedical applications are hyaluronan, chondroitin sulfate, heparin and alginate, which are all polyanions and chitosan, a polycation (Figure 8.1). These polysaccharides are formed by dimeric sugar molecules. Usually one sugar is a uronic acid (either D-glucuronic acid or L-iduronic acid) and the other is either N-acetylglucosamine or N-acetylgalactosamine. One or both of the sugars contain one or two sulfate residues. Thus each polysaccharide (also called glycosaminoglycan) chain bears many negative charges, either carboxylic or sulphate groups. Chitosan (CHI) is a linear polysaccharide containing two β-1-4 linked sugar residues, N-acetyl-D-glucose amine and D-glucosamine. It is obtained by partial N-deacetylation of chitin from crustacean shells, chitin being the 231 © 2008, Woodhead Publishing Limited

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Natural-based polymers for biomedical applications Chitosan OH NH2 O

HO O

OH

NH2 O HO

O

HO O

O

HO

NH2

NH2

n

OH

OH

O

O

O

Hyaluronan OH

OH

O

HO O O

O

O HO

OH

O NH

O

O

OH HO O O

NH

O HO

O n

OH

O

O OH

Chondroitin sulfate O +

OH

O

Na–O S HO O

O

O HO

O

O

O

O NH

OH

OH

HO O O

O

NH O HO

O

O

O O

OH O

n

S

O– Na+

O Heparin O HO OH

O

O HO

O

S O

OH

O

O

HO

S

S NH

O

R′ O

O O

HO

S

O O

O

O O

O

HO O

NH

O

O

O

n

O

OH

S O

O HO

O

OH

S O

OH

Alginic acid

O HO

OH OH O

O

O HO O

OH

OH

OH OH O

O

OH O

O HO O

HO O

OH O O

n

8.1 Schematic of the molecular structures of the natural polysaccharides that will be evoked in the chapter: chitosan, hyaluronan, chonohoitin sulphate, heparin, alginate.

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second most abundant naturally occurring polysaccharide.17 Chitosan is the only natural polycation. Hyaluronan (HA) is also a linear polysaccharide constituted of alternated N-acetyl-β-D-glucosamine and β-D-glucuronic acid residues. Hyaluronan possesses lubricating functions in the cartilage, participates in the control of tissue hydration, water transport, and in the inflammatory response after a trauma. 16 These polysaccharides are biocompatible, non-toxic and biodegradable by enzymatic hydrolysis with chitosanase,18 α-amylase19 lysozyme, and hyaluronidase.20 Both have already been widely used in a variety of biomedical applications, such as tissue engineering14,21,22 controlled drug release or capsule formation.17,23 Both polysaccharides have a relatively high intrinsic chain stiffness, with persistence length of ~6 nm for hyaluronan24 and of ~6–12 nm for chitosan.25,26 Chitosan and hyaluronan can be easily chemically modified27–29 and coupled to various molecules such as cell-targeted prodrugs,30 carbohydrates,31 which could be released during film hydrolysis. The structure of chondroitin sulfate (CS) is close to that of hyaluronan except that it bears a sulfate group on the N-acetyl-β-D-glucosamine. Chondroitin sulfate is present in the interphotoreceptor matrix and is used as a component of skin substitutes.32 It can also serve for encapsulation and subsequent delivery of drugs in the treatment of colon-based diseases.33 Heparin (HEP) is also a linear anionic polysaccharide chain that is typically heterogeneously sulfated on alternating L-iduronic acid and D-glucosamino sugars. It is highly charged and can be considered as a strong polyelectrolyte, contrary to all the other polysaccharides. Heparin is well-known to show anticoagulant activity. Alginate (ALG) or sodium alginate is the sodium form of alginic acid (Figure 8.1). Alginates are naturally occurring polysaccharides that are found in algae. Alginates are copolymers containing mannuronic acid (M) and guluronic acid (G) monomeric subunits of varying amounts and distribution along the polymer backbone. Its form as a gum, when extracted from the cell walls of brown algae, is used by the foods industry to increase viscosity and as an emulsifier. It is also used in indigestion tablets and the preparation of dental impressions. Also, due to alginate’s biocompatibility and simple gelation with divalent cations, it is widely used for cell immobilization and encapsulation. In particular, poly(L-lysine) (PLL) and alginate is a polymeric system that has been widely used for the coating of microcapsules.6 This polymer system is capable of forming complex coacervates at physiological conditions, has already demonstrated a degree of bioinertness and is capable of forming very thick coatings that can be generated around microcapsules. Finally, collagen (COL) is a natural polymer, which is a major structural protein in tissues. It exists in different forms with type I being the most common. Its tertiary structure forms triple helixes, and this entity is physicochemically stable in solution. Its quaternary structures consist of the collagen

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fibrils and fiber, and they are stable in the solid form. Collagen provides the necessary environment for cell attachment and is a ligand for certain cell surface receptors (such as integrins). In the present chapter, we will focus on natural based films made of CHI, HA, ALG, CS, HEP or COL, for at least one of their components. The films made from synthetic polyelectrolytes or from poly(aminoacids) that contain natural materials such as proteins will not be considered here.

8.2

Physico-chemical properties

8.2.1

Film growth: Linear versus exponential

The first investigated polyelectrolyte systems described by Decher and coworkers4 exhibited a linear growth of both the mass and the thickness of the films with the number of deposition steps. Poly(styrene sulfonate)/ poly(allylamine hydrochloride) is one of the most prominent examples of a linearly growing system.34–39 These films present a stratified structure, each polyelectrolyte layer interpenetrating only its neighbouring ones. The growth mechanism involves mainly electrostatic interactions between the polyelectrolytes from the solution and the polyelectrolytes of opposite charge forming the outer layer of the film. Each new polyelectrolyte deposition leads to a charge overcompensation that is the actual motor for the film growth and to a change in the zeta potential.34 More recently, using polysaccharides and polypeptides, Elbert and coworkers6 and Picart and co-workers40,41 described a new type of polyelectrolyte multilayer which is characterized by an exponential growth of both the mass and the thickness of the film with the number of deposition steps. PLL/ALG6 and PLL/HA40,41 were the first reported examples. CHI/HA9 and PLL/CS42 are other examples. Whereas the typical thickness of a linearly growing film consisting of 20 layer pairs is of the order of 100 nm, the thickness of exponentially growing films, in a physiological medium, can reach 4 µm or more after the deposition of a similar number of layers (Figure 8.2). We reported that the construction of poly(L-lysine)/HA films took place over two build-up regimes. One consists of the formation of isolated islands of the PEM that grows to a continuous film, whereby the second regime sets in, characterized by an exponential increase of mass with the number of added layers. Other exponentially growing films have been reported.43,44 Two explanations for these exponential growth mechanism have been proposed: one relies on the diffusion of polyelectrolyte ‘in’ and ‘out’ of the film during each ‘bilayer’ step41,43 while the second one relies on the increase in film surface roughness as the film builds up.35,45 However, no change in surface roughness was observed for the exponentially growing films made of polypeptides.40,43,46 A deep investigation of the PLL/HA system allowed us to better understand the processes underlying such a growth mechanism.41 © 2008, Woodhead Publishing Limited

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2003 1803 1603 1403

–∆f / v

1203 1003 803 603 403 203 0 1

2

3

4

5

6

7

8

5

6

7

8

(a)

250

D (at 15 MHz)

200

150

100

50

0 1

2

3

4 (b)

8.2 Natural-based multilayer film growth. Exponential growth of (PLL/ HA) (䊉) and of (CHI/HA) films (䊊) followed by using a quartz crystal microbalance. Different parameters are represented during the alternation of PLL (resp. CHI) and HA layers on SiO2 crystal: (a) frequency shift (–∆f/ν) measured at 15 MHz, (b) viscous dissipation measured at 15 MHz, (c) thickness deduced from the fit of the QCM data at the four frequencies and dissipations.

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Natural-based polymers for biomedical applications 350

300

Thickness (nm)

250

200

150

100

50

0 1

2

3 4 5 Number of layer pairs (c)

6

7

8

8.2 (Continued)

For this investigated system, the growth mechanism relies on the diffusion ‘in’ and ‘out’ of the whole structure during each ‘bilayer’ deposition step of one type of the polyelectrolytes constituting the films.41,43 The diffusion of PLL and CHI could be visualized by confocal laser scanning microscopy (CLSM) for the (PLL/HA) and (CHI/HA) films using fluorescently labeled polyelectrolytes (respectively PLLFITC and CHIFITC) (Figure 8.3). Diffusion of PLL was also observed by CLSM for PLL/CSA films.42 Most, but not all, of the reported exponentially growing films contain PLL or CHI as polycation. It was also evidenced that a polyanion/polycation system that grows exponentially under certain conditions can become linearly growing when the deposition conditions are changed. This is particularly the case when the salt concentration is varied from low, corresponding to a linear growth, to high, corresponding to an exponential growth. This was evidenced for CHI/ DEX films by Serizawa et al.7 and for CHI/HA films by Richert et al.9 The simplest explanation is that, by reducing the salt concentration of the polyelectrolyte solutions during the buildup, the films become thinner (for a given number of deposition steps) and more dense, thereby hindering polyelectrolyte diffusion into the film. Interestingly, films containing collagen were found to grow linearly.47,48 It was also shown that, for the linearly growing films like those containing collagen, vertical diffusion of the collagen of the film did not occur and collagen adsorbed on top of the film.47 Table 8.1 summarizes all the different systems investigated, the buildup conditions, and the type of growth (linear or exponential).

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(a)

(b)

8.3 Confocal images of (PLL/HA)24-PLLFITC (a) and (CHI/HA)24-CHIFITC films (b). Vertical sections through the films containing the labeled polycation are shown. The glass substrate (bottom of the chamber) is indicated with a white line. The image sizes are respectively 45 µm × 8 µm and 45 µm × 12 µm. Green fluorescence (corresponding to PLLFITC and CHIFITC) is visible over the total film thickness, i.e. ~ 4 µm for (PLL/HA)24 films and ~ 6 µm for (CHI/HA)24 films.

Table 8.1 Studies involving natural based multilayer films, either with poly(L-lysine), chitosan, or collagen as polycations. Experimental conditions and type of growth are given Study

PLL as polycation PLL/Alginate Elbert et al. (6)

Conditions

Type of growth

PBS

Exponential

Picart et al. (40)

0.15 M NaCl pH 6.5

Exponential

PLL/CSA

Tezcaner et al. (42)

0.15 M NaCl pH 6

Exponential

PLL/Heparin

Boulmedais et al. (112) 0.15 M Nacl + Hepes buffer, pH 7.4

PLL/HA

CHI as polycation CHI/HA Richert et al. (9) Kujawa et al. (113)

Exponential

0.15 M NaCl pH 5

Exponential

CHI/Heparin

Fu et al. (96)

0.15 M NaCl pH 3 to 3.8

Linear

CHI/Mucin

Svensson et al., (114)

Acetic acid, no salt pH=4

Linear

CHI/Dextran sulfate CHI/HEP

Serizawa et al. (7)

NaCl at different concentrations

Linear for NaCl < 0.5 M Exponential for 0.5M et 1M NaCl

COLLAGEN as polycation COL/HA Zhang et al. (47) COL/HA

Johansson et al. (48)

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0.05M NaCl, pH = 4

Linear

0.1 M Acetate buffer pH 4

Linear

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8.2.2

Natural-based polymers for biomedical applications

Film hydration and swellability: Sensitivity to external parameters such as pH and ionic strength

Film hydration can be estimated by measuring the film refractive index using techniques such as optical waveguide lightmode spectroscopy40 or ellipsometry (by measuring respectively dry and hydrated film thickness).49 Refractive index of synthetic polyelectrolyte multilayer films, such as poly(styrenesulfonate)/poly(allylamine hydrochloride) (PSS/PAH) films, was measured in situ by OWLS and a value of 1.5 was estimated in physiological conditions.39 This indicates that such films are relatively dense and contain only around 25% of water (a simple approximation of the water content is based on the following formula : nPEM = 1.3340×a + (1–a)×1.56, 1.334 being the refractive index of a 0.15 M NaCl solution, 1.56 being the refractive index of a pure polymer film,49 and a being the fraction of water). Other studies were realized with PLL, poly(D-lysine) (PDL), or even chitosan as polycation in combination with polyanions such as gelatin50, poly(L-glutamic) acid (PGA),51 or hyaluronan.9 In general, films made of polypeptides and polysaccharides in comparable ionic strength conditions are more hydrated than films made of synthetic polyelectrolytes such as PSS/PAH. This observation is based on refractive indices that are ≈1.36–1.38 for polysaccharide films9,40 and ≈1.42 for PGA/PLL films,43 which would correspond to water contents ranging respectively from 95% to 60%. This refractive index for (PLL/HA) films is of the same order of magnitude as that found by ellipsometry49 for wet films, (1.35) and has to be compared to the refractive index for dried films. (1.56) This indicates that the film swells by about 830% (initial conditions for film assembly were pH 9 and 0.1 M NaCl). The high swelling capacities of the polysaccharides, and in particular for hyaluronan,52 renders the buildup of much thicker films possible, up to several hundreds of nanometers9, or even several micrometers after deposition of 20 to 30 layer pairs.41 These polysaccharide based films were often, if not always, found to be extremely cell resistant,6,9,53 except when the films were rigidified by covalent cross-linking.53 Therefore, a trend that seems to emerge from all these cell lineages and primary cell studies is that nanometer thin and dense films formed by few layer pairs are more favorable for cellular adhesion than thick and highly hydrated films. A detailed study of the hydration and swelling properties of (PLL/HA) films indicates that the most important parameters are: (a) the assembly pH (which can be varied from 5 to 9 for these particular films) and ionic strength; (b) the swelling medium pH and ionic strength.49 Thus, depending on the combination of these parameters, very different film properties can be achieved. Polysaccharides like HA have, in particular, the ability to adopt secondary structures and can exhibit H-bonded helical

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conformations accompanied by chain stiffening, when their charge fraction is low (i.e. at low pH close to the pKa). They can also exhibit hydrophobic interactions,16 which are influenced by ionic strength. Very interestingly, the measurement of the pKa of the polyelectrolytes in the film demonstrates that both PLL and HA experience a significant shift in their pKa(apparent) values upon adsorption, compared to the accepted values (in dilute solution) of 9.36 ± 0.08 and 3.08 ± 0.03 respectively in the presence of 1.0 mM NaCl. The pKa(apparent) values of both PLL and HA remained relatively constant after the first 3–4 deposited layers (at pH 7, it is 4.85 for HA and 6.8 for PLL). Such decrease in the acid strength of HA and base strength of PLL is similar to that reported for other polyelectrolyte pairs.54 It has been previously speculated and experimentally shown that the charge on the multilayer film surface strongly influences the acid-base equilibria of adsorbing polyelectrolyte chains.55 According to Barrett et al., for PLL/HA multilayer films, the overall trend in the pKa(apparent) shifts upon adsorption, in comparison to the dilute solution values, are influenced by the ability of both of these polymers to adopt some degree of secondary conformational order with changes in the local pH and ionic strength environment.49,56,57,58 In the intermediate pH range, HA is known to have some degree of chain stiffening in solution due to local hydrogen bonded helical regions, whereas PLL chains are reported to experience a random coil to α-helix transition at pH = 10.5.59 The same authors also investigated the swelling of PAH/HA films and found that these films exhibit a high dependence of swelling on the assembly solution pH. The swelling ratio varied between two at physiological pH of 7 to more than eight at very acidic pH of 2 and was more pronounced than at basic pH of 10 (swelling ratio about five).

8.2.3

Stability in physiological medium

Although, in principle, multilayer films can be built under very different conditions in terms of pH and ionic strength, the final suspending medium may depend on the foreseen application. In particular, when cell culture studies or deposition on biomaterial surfaces are foreseen, it is then necessary that the films are stable in culture medium and in physiological conditions. These requirements may greatly limit the range of possible buildup conditions due to stability constraints. On the other hand, if the films are to be used for a subsequent release of a film component itself or of a bioactive molecule (see below), then, stability is not a matter or at least, is not as important as in the first case. It stems from the aforementioned properties of the natural-based multilayer films (weak electrostatic charge, high hydration and swellability, secondary interactions) that these films can be subject to stability problems. This is particularly true when the films are built in a medium which has a different

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pH and/or ionic strength from a physiological medium (ionic strength of about 0.15 M NaCl, neutral pH). Then, the films are subject to stresses upon medium change and can potentially be disrupted due to too high internal stresses. Typical cases are films built at acidic pH like COL/HA films or CHI/HA films, for which COL and CHI are polycationic at acidic pH (4 for COL and less that ~5.5 for CHI). Johansson et al. found that COL/HA films are not stable when the pH is raised from 4 to 7.48 This could be explained by the protonation/deprotonation process for the polyelectrolytes involved in the interaction. At pH 4.0, most acid functionalities are protonated, whereas they are deprotonated at pH 7. Regarding collagen, the number of negatively charged acids on collagen approaches the number of protonated amines or the isoelectric point. The protonation/deprotonation processes induces the changes in the three-dimensional structure of the polyelectrolytes, which affects the electrostatic forces that existed between the polyelectrolyte layers. This dissolution was found to be irreversible. Regarding CHI/HA films, we found that the stability depends on the molecular weight of the chitosan: whereas films built with high molecular weight chitosan are stable in physiological medium,9 films built with chitosan oligosaccharides (MW 5000 g/mol) exhibit a change in structure when introduced into the culture medium.60 We evidenced that this change in structure was mostly due to the presence of divalent ions (Ca2+, Mg2+ in the culture medium) and not to the change in pH. In fact, divalent ions are known to complex chitosan61 and also alginate.62 These observations are not only valid for polysaccharide multilayer films but for other sensitive films like hydrogen-bonded films built at very low pH63 and PLL/PGA films built at low pH.64 Even when films are not built in acidic or basic conditions, they may be subjected to dissolution in a physiological medium. This was observed for films containing PEI as polyanion and a mixture of heparin and acid fibroblast growth factors whose degradation could be observed in PBS at 37°C.65 On the contrary, films built with basic fibroblast growth factor and chondroitin sulfate66 were stable in PBS. However, it is difficult to establish a common rule and each type of film needs to be tested. It must also be noticed that the presence of cells, which are able to exert strong stress on their matrix,67 can also affect the film stability. We will see below that such problem of stability in physiological medium and of mechanical resistance can be overcome by cross-linking the films.

8.3

Different types of natural-based multilayer films for different applications

8.3.1

Supported films

Most studies of natural-based multilayer films are performed on planar substrates. These films are called ‘supported films’. Depending on the © 2008, Woodhead Publishing Limited

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experimental technique for probing the film buildup or investigating the cell/ film interactions, the material surface is most often silicon, gold, or bare glass. Atomic force microscopy, CLSM observations and UV-visible spectroscopy are commonly performed on glass or quartz slides.41,43 Quartz crystal microbalance experiments make use of SiO2 or gold coated crystals. However, it is important to note that a great advantage offered by PEM films is their ability to coat any type of material with any shape. Thus, PEM films have recently been deposited onto stainless steel,68 polydimethysiloxane (PDMS),50 vascular stents made of NiTi,10 and onto biodegradable poly(Llactic) acid matrices.69 The geometry was not necessarily planar but also curved or spherical, as for titanium beads,70 or polystyrene and glass microspheres. We observed by scanning electron microscopy the deposition of (PLL/HA)24 films on polyethylene terephthalate filaments, on NiTi and on stainless steel surfaces (Figure 8.4). It is clearly visible that the film is smoothing the initially rough surface and it is entirely covering the surface. The side view image of film-coated NiTi surface by CLSM also shows that the film homogeneously covers the materials (data not shown).

8.3.2

Capsules (drug release)

In the case of films built on particles, the particle core can also be subsequently removed to form hollow capsules.8 The capsules offer broad perspectives in drug delivery.71,72 The main advantages of polyelectrolyte capsules are their large versatility and modularity according to the materials and conditions used for their preparation. To date, the most studied capsules are of poly(allylamine)/poly(styrenesulfonate) (PAH/PSS). Several properties such as permeability73–76 and stability against environmental alterations such as pH and temperature77,78 of the capsules have been investigated. Owing to the potential applications of the capsules in biology, the use of natural polysaccharides and derivatives, which have the advantages of biocompatibility, biodegradability and in some cases, bioactivity, has emerged to prepare LbL capsules. Berth et al. examined the buildup and permeability properties of chitosan/chitosan sulfate capsules prepared by deposition of the films on a melamine formaldehyde latex template.79 Depending on the pH and ionic strength of the suspending medium, the capsules exhibit different aspects (core or shell labeled, or both). Zhang et al. made use of the LbL technique to prepare single component hollow capsules made of chitosan.80 Toward this end, they built a film containing chitosan and poly(acrylic acid) (PAA), cross-linked it with glutaraldehyde and subsequently removed PAA by placing the capsule in a carbonate buffer at pH 9. The monodispersity of the single component capsules was proven by dynamic light scattering measurements. Recently, dextran sulfate was associated with chitosan to prepare enzymeresponsive biodegradable hollow capsules.81 The capsules were sensitive to

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(a)

(b)

(c)

(d)

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8.4 Natural-based film coated biomedical materials. Scanning electron microscope images of bare materials (a, b, c) and of (PLL/ HA)24 film coated materials (d, e, f): (a, d) polyethylene terephthalate; (b, eE) stainless steel; (c, f) Nickel-titanium alloy. (scale bar is 20 µm for a, 50 µm for d, and 10 µm for all other images).

enzymatic degradation by chitosanase and could release albumin-FITC that was entrapped in the capsule core. Other polysaccharides like alginate and carboxymethyl cellulose76,79–81 have also been introduced as polyanions for the fabrication of capsules. Up to now, however, the use of hyaluronan, probably the most hydrated of all the polysaccharides, for the fabrication of hollow capsules, remains unexplored. Compared to the planar films made from natural polysaccharides, little is known about capsules with polysaccharide nanoshells. The recent development of more complex synthetic double wall capsules (or ‘shell-in-shell’ capsules)82 will probably be applied, in the next few years, to more biomimetic components. The development of new types

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of capsule made from polysaccharides will thus constitute a new challenge for the next few years.

8.3.3

Membranes

When the polyelectrolyte multilayer films are detached from the surface, they give rise to self-supported membranes.83 Although many studies have focused on the preparation and characterization of membranes, the membrane constituents are generally synthetic polyelectrolytes such as PSS, PAH or poly(diallyldimethylammonium chloride) (PDADMAC).84 Few have used polypeptides as film constituents85 and only few studies report the preparation of a membrane containing polysaccharides.86,87 Miller and Bruening prepared different membranes whose swelling and transport properties were investigated by ellipsometry and nanofiltration (NF) rejections and diffusion dialysis fluxes.86 They found that hyaluronic acid (HA)/chitosan films swell four times more than poly(styrene sulfonate) (PSS)/poly(allylamine hydrochloride) coatings, and in NF experiments, the HA/chitosan membranes permit a 250fold greater fractional passage of sucrose. In general, films prepared from polyelectrolytes with a high charge density showed low swelling and slow solute transport, presumably because of a high degree of ionic cross-linking. These results are in agreement with the previous findings on the high swellability of polysaccharide multilayer films40,49 and confirm that swellability is related to permeability. Recently, Kotov et al. prepared composite membranes made of chitosan and of montmorrillonite (MTM) with a high loading of MTM comparable to that in the natural nacre (~80%). In contrast to the theoretical predictions, these membranes exhibited lower strength and stiffness than those of poly(diallydimethylammonium) (PDDA)/MTM. The authors concluded that CHI, although a much stronger polysaccharide polycation than PDDA, lacks the flexibility necessary for strong adhesion between the organic matrix and MTM platelets.88 Lavalle et al. investigated the formation of (PLL/HA) membranes.87 As (PLL/HA) films are soft and sensitive due to their high hydration, the common protocol of detachment that consists of dipping the film-coated polystyrene substrate into tetrahydrofuran (THF), leads to the formation of micrometric holes (several tens of micrometers) in the membrane. The authors had to develop an alternative strategy based on: (a) the increase in film mechanical properties by cross-linking via a carbodiimide; (b) the detachment of the silica surface by dipping the film in a 0.1 M NaOH solution (pH 13). Using this protocol, they could obtain a homogeneous and smooth membrane and could indeed functionalize it by a model enzyme, alkaline phosphatase.

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8.4

Bioactivity, cell adhesion, and biodegradability properties

8.4.1

Bioactivity based on the film components (e.g. heparin)

The natural polyelectrolytes can give specific properties to multilayers due to their intrinsic properties. For instance, chitosan anti-bacterial properties have received considerable attention in recent years.17 The exact mechanism of antibacterial action of chitosan is still unknown but various mechanisms have been proposed: interaction between positively charged chitosan molecules and negatively charged microbial membranes leads to the leakage of intracellular constituents. Heparin, with its anti-thrombogenicity and strong hydrophilicity, prevents adhesion of bacterial cells and is an excellent candidate for anti-adhesive coatings. Chitosan/dextran films were found to exhibit anti-coagulant properties only when dextran is the outermost layer of the film and when the films are built in 0.5 M NaCl or 1M NaCl. On the other hand, chitosan/heparin films built in 1M NaCl also exhibited strong anticoagulant activity whatever the outermost surface of the film.7 Thus, such multilayer films have good potential for the surface modification of medical implants in contact with blood. The thromboresistance of a (CHI/ HA)4 coated NiTi substrate was also evidenced by Thierry et al.10 These films were found to significantly reduce platelet adhesion, by 38%, after one hour exposure to platelet rich plasma. On the contrary, the adhesion of polymorphonuclear neutrophils increased slightly on the coated surface, compared to bare metal.

8.4.2

Bioactivity based on the insertion of bioactive molecules in natural based multilayer films

Beside the intrinsic properties of the polysaccharides that constitute the film, it is possible to benefit from the high swelling properties of these films and from their large thickness for using them as reservoirs for drugs or bioactive molecules. It is precisely because these films have a low degree of ionic cross-links and a large porosity that they can be employed as reservoirs. Therefore, not only can small molecules be loaded in the films but also proteins like myoglobin, which was found to diffuse within CHI/HA films.89 Thierry et al. found that the incorporation of sodium nitroprusside, a nitrous oxide donor that is widely used clinically to reduce blood pressure, within the (CHI/HA) coating further decreased platelet adhesion by 40%. The reservoir capacity of thick films was nicely evidenced by Vodouhe et al.90 Using (PLL/HA) film as a matrix, they evidenced, using CLSM, that paclitaxel Green 488 molecules diffuse through the whole (PLL/HA)60 film section

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and that the fluorescence is homogeneously distributed over the whole film thickness. They successfully increased the amount of drug uptake by increasing the paclitaxel solution concentration. They found that the effective concentration in the film was from 20 to 50 times greater than the initial solution concentration. For instance, when the solution concentration was 10 µg/mL, the effective concentration in the film was 500 µg/mL. Using this method, the drug content in PLL/HA films can be finely tuned in a large concentration range. A similar strategy was employed by Schneider et al., who loaded cross-linked (PLL/ HA) films with the anti-inflammatory drug sodium diclofenac and with paclitaxel. The amount of drug loaded could be tuned by varying the film thickness.91 The effect of paclitaxel, loaded in the cross-linked (PLL/HA) films, could be clearly seen over the three days culture period (Figure 8.5). After three days in contact with the bioactive films, less than 10% of the cells were still alive.92 Larger molecules like adenovirus (Ad) particles or even proteins like growth factors can be adsorbed onto or embedded in natural-based films.93 The Ad particles, which are 70 nm in diameter, were found to adsorb on (PSS/PAH) film surface and to be partially embedded in the multilayer films. They were even found to diffuse within (PLL/HA) films. The bioactivity of

100

**

AP activity (%)

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60 *** 40

20

***

0 24 H

48 H

72 H

8.5 Acid phosphatase (AP) activity for HT29 cells cultured on crosslinked (PLL/HA)12 films loaded (cross-hatched) or not (black) with paclitaxel, after time periods of 24H, 48H and 72H in culture. The error bars represent the standard deviation. The value of 100% has been arbitrarily set at 100% for CL films at each time period (** p<0.01; *** p<0.001 versus controls, which are the CL films at time 24H, 48H, and 72H respectively). (Adapted from ref 92.)

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the particles was preserved when adsorbed in (PLL/HA) (18% remained infectious) and (PLL/PGA) films (24% remained infectious). Interestingly, whereas the Ad particles enveloped by (PLL/PGA) and (PSS/PAH) films (two to six layer pairs) remain inaccessible to cells, in PLL/HA and CHI/HA films, the overlay Ad with the same number of layers is responsible for a progressive and less-important decrease in cell transduction. The authors postulate that Ad diffusion through the multilayers make it more accessible for cellular uptake and/or more available for cell infection. The maximal cell-transduction efficacy could be achieved when Ad was well embedded and overlaid in two levels in PLL/HA and CHI/HA films. Shen and co-workers introduced a new class of bioactive films using direct growth factors (acidic or basic fibroblasts growth factors, aFGF or bFGF respectively) as building blocks, either mixed with heparin and deposited alternately with PEI, or directly used as polycation and deposited with chondroitin sulfate A.65,66 An enhanced secretion of collagen type I and interleukin 6 (IL-6) by fibroblasts seeded on the five layer pairs of (aFGF/ heparin)/PEI was also observed by immuno-histochemistry. When bFGF was directly built in multilayer films with CSA, the films containing bFGF had an improved bioactivity. In vitro incubation of the CS/bFGF multilayers in PBS showed that about 30% of the incorporated bFGF was released within eight days. The fact that the growth factors retained their biological activity is very interesting for biomedical applications. Another way to deliver a drug is to couple it to the polyelectrolyte via a hydrolysable linkage. This so-called prodrug approach was employed by Thierry et al. for loading and subsequently releasing paclitaxel from CHI/ HA-paclitaxel films.94

8.4.3

Cell and bacterial adhesion or anti-adhesion

Most of the natural-based multilayer films exhibit anti-adhesive properties, presumably due to their high hydration and softness. This is valid for (CHI/ HA) films on top of which adhesion of primary chondrocytes was extremely low (less than 4% of the control) for both CHI and HA ending films,9 and for (PLL/HA) films which are very ‘cell resistant’ toward chondrosarcoma cell adhesion.53,95 Also, PLL/alginate films were found to be bioinert and to resist fibroblast adhesion.6 Surprisingly, however, photoreceptor cells could adhere and remain viable when plated onto PLL/CSA and PLL/HA films.42 Beside cell adhesion, bacterial adhesion (E. coli Gram-negative strain) was also investigated on certain types of natural based multilayer films containing chitosan and/or heparin. (CHI/HA)10 films (built in 0.15 M NaCl) are highly resistant to bacterial adhesion and lead to a ≈80% decrease in bacterial adhesion as compared to bare glass.9 On the other hand, (CHI/ HA)20 films built in 10–2 M NaCl were less resistant to bacterial adhesion

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(40% less than the control on the CHI ending films and 20% less than the on the HA-ending films). The observed differences were explained by the lower thickness of the (CHI/HA) films built in 10–2 M NaCl (120 nm as compared to 300 nm for those built in 0.15 M NaCl) and also probably by an increased film rigidity. The (CHI/HA)10 films built in the presence of 0.15M NaCl could thus be used as anti-microbial coatings for biomaterial surfaces. In another study, multilayer films made of chitosan and heparin were found to kill bacteria adhered to the surface. E.coli initial adhesion was also greatly decreased on the multilayer film.96 The assembly pH was found to be an important parameter in the design of efficient anti-adhesive and antibacterial films.

8.4.4

Biodegradability properties: Studies in vitro and in vivo

The biodegradability of ultrathin polymer films coated on material surfaces is one of the most important requirements for biomedical applications of these polymers. Due to their intrinsic properties and to the presence of specific enzymes in vivo, natural-based multilayer films can be degraded in the presence of enzymes and thus release their content. Serizawa et al. were the first to demonstrate the alternating enzymatic hydrolysis of an LbL assembly formed from chitosan and dextran.18 Chitosanase, an enzyme which hydrolyzes chitosan, was applied in this process. More recently, we showed that CHI/ HA films can be degraded by enzymes such as hyaluronidase, lysozyme and α-amylase, which are present in saliva and plasma.97 The biodegradability of the films could be tuned by varying the extent of film cross-linking60, using a carbodiimide at various concentrations. In addition, phagocytic cells such as THP-1 macrophages can degrade (PLL/HA) native films whereas the cross-linked films are not degraded over the same time period.98 Figure 8.6 show the biodegradability of (PLL/HA) and (CHI/HA) films in contact with THP-1 macrophages. (a) (b)

8.6 Biodegradability of polysaccharide multilayer films. CLSM study of the biodegradability of a native (PLL/HA)24-PLLFITC (a) and of a cross-linked (CHI/HA)24-CHI-FITC film (cross-linked at 5 mg/mL EDC concentration) (b) both having been in contact with THP-1 macrophages for 24 hours at 37°C. Cross-sections of the films are shown (image size is 230 µm × 12 µm for (a) and 230 µm × 22 µm for (b)). Arrows indicate the visible degradation sites. (Adapted from refs 60 and 98.)

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Biodegradability of the films in vivo remains barely explored. In our group, we investigated the biodegradation of CHI/HA films, cross-linked to different extents, in two different locations: the rat oral cavity (for possible applications in the dental field)99 and the mouse peritoneal cavity (for investigating possible tissue implantations).60 We observed that the native films were very rapidly degraded due to the presence of saliva enzymes and probably due to strong mechanical stresses exerted on the films. On the other hand, about 60% of the cross-linked films remained intact after three days in the oral cavity. When implanted in the peritoneal cavity, the (CHI/HA) films induce an inflammatory response that directly depends on the extent of cross-linking: the more cross-linked the film is, the more macrophages it attracted. However, the highly cross-linked films were also more resistant to degradation. Thus, it seemed that there is an optimal cross-linking for, at the same time, increasing resistance and mechanical stability in vivo but not inducing a too high inflammatory response.60

8.5

Modulation of film mechanical properties

Characterizing and modulating the film mechanical properties has become a challenge of the last few years. In particular, for natural-based multilayer films which are rather soft as compared to films made of synthetic polyelectrolytes, it is important to ensure that they will be stable in different applications and, in particular, when they are submitted to different stresses (like shear stress). Mechanical properties of the films can be modulated in different ways.

8.5.1

Mechanical properties based on the structure of the polyelectrolyte

Schoeler et al. investigated the buildup of films containing PAH as polycation and two ionic polysaccharides, iota and lambda-carrageenan, of similar composition but different conformations.100 Iota carrageenan is in a helical conformation at room temperature whereas lambda carrageenan is in a random coil conformation. The mechanical properties of these two film types were found to be very different, which suggests that the structure of the film can strongly influence its mechanical properties. Also, the presence of different sugar molecules (like lactose or mannose) is sufficient to create noticeable differences in film stiffness and to modulate cell adhesion.101

8.5.2

Modulation of mechanical properties by film cross-linking

For increasing the film mechanical properties, different strategies have been proposed and some of them have been applied to natural-based films. One of © 2008, Woodhead Publishing Limited

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the strategies is to incorporate nanocolloids in the films.102,103 Although not directly applied to natural-based films, one can observe that the CHI/HA and PLL/HA films containing adenovirus particles (stiff particles of 70 nm in diameter) are a good surface for cell adhesion, whereas films that do not contain particles are non-adhesive.9,53 Another strategy consists of ‘capping’ the ‘soft’ films with a layer of a stiffer polyelectroyte, like PSS. This was applied by Vodouhe et al. who observed that initially non-adhesive PLL/HA films could be rendered adhesive by the addition of a single PSS layer.90 Presumably, there is an exchange between the HA molecules outside the films and the incoming PSS molecules. Another possibility would be to prepare films containing a mixture of polyelectrolytes, such as a mixture of hyaluronan and PSS104 in combination with PLL, one of the mixture components being a strong polyelectrolyte. Finally, it is possible to create covalent cross-links within the films by making use of the existing charged groups in the films. In our group, we developed a protocol based on carbodiimide chemistry for cross-linking carboxyl groups with amine groups and thereby creating covalent amide bonds.53 This protocol was applied to several types of polyelectrolyte pairs, including PLL/HA and CHI/HA films.60,105 The effective cross-linking was checked by Fourier Transform Infrared Spectroscopy in attenuated total reflection mode (Figure 8.7). Several peaks and in particular the carboxylic peaks (at 1606 and 1412 cm–1) and the polysaccharide C-O stretching at 1082 cm–1 and 1032 cm–1 decreased (the polysaccharide peaks were only clearly visible for CHI/HA films) whereas at the same time the intensity of other bands increased. This is the case for the amide I and II bands (in the 1630–1700 cm–1 region and in the 1440–1500 cm–1 region respectively) and, in the case of CHI/HA films, for the C-O ester band in the 1180–1260 cm– 1 region106 and for the C=O ester band at 1740 cm–1. The decrease of the carboxylic peaks of HA and the concomitant increase in the amide bands proves the reaction between the corresponding chemical groups and the ammonium groups of PLL and of CHI. In the case of CHI/HA films, the disappearance of the characteristic saccharide peaks and the appearance of ester bands at around 1240 cm–1 and at 1740 cm–1 suggest the formation of ester bonds. Such bonds involve hydroxyl groups of the polysaccharide and carboxylic groups or acid anhydride formed by the reaction between two carboxylic groups.107 Very interestingly, we found two important consequences of the cross-linking, which are probably related: (a) a change in the film mechanical properties; and (b) a drastic change in the film adhesive properties. The film’s mechanical properties have been investigated by means of AFM nano-indentation experiments.105,108,109 We found that the Young’s modulus (E0), which represents the mechanical stiffness of the films, is ~5 kPa for native films whereas it can be varied over two orders of magnitude from 5 to 500 kPa for cross-linked films (depending on the (carbodiimide) cross-

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Natural-based polymers for biomedical applications 0.3 Ester Saccharide peaks C–O

0.2

COO– 0.1

Ester C=O COO– Amide I

Amide II

Absorbance (AU)

0.0 –0.1 –0.2 –0.3 –0.4 –0.5 –0.6 900

1000

1100

1200

1300 1400 1500 Frequency (cm–1)

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1700

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8.7 Film cross-linking followed by Fourier Transform Infrared Spectroscopy in Attenuated Total Reflection mode. Differences between the ATR-FTIR spectra taken before and after the crosslinking procedure (e.g. after the final rinsing step) are represented for two different film types: a (PLL/HA)8 film (dotted line) and a (CHI/HA)9 film (continuous line). Spectra are shown over the frequency range from 900 cm–1 to 1800 cm–1 (Adapted from refs 53 and 98.)

linker concentration). This was the first time that the Young’s modulus of such thin films could be systematically varied and tuned. With respect to the cell behavior on top of cross-linked films, we have already shown that the ‘switch’ from non-adhesive to adhesive upon crosslinking is a common property for many films types, including PLL/HA,53,98,105 CHI/HA,9,109 and PLL/PGA110 and was observed for a large variety of cell types (chondrosarcomas,111 chondrocytes,98 osteoblasts,110 neurons,98 smooth muscle cells105). We also observed that proliferation is enhanced on the cross-linked films as compared to the native ones or as compared to films grafted with an adhesive peptide (a 15 amino-acid peptide containing the RGD sequence).110 We are now investigating cell differentiation onto these films and have evidenced a strong influence of film mechanical properties on differentiation (data not shown).

8.6

Future trends

The physico-chemical properties of the natural based polyelectrolyte multilayers need to be better understood. In particular, their porosity and the diffusion of

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molecules of various sizes (from small molecules to proteins) has to be understood and controlled. The role of chain stiffness, which is an important parameter for polysaccharide molecules, could be studied. Similarly, hydrophobicity, which could play an important role in the insertion of amphiphilic growth factors or hydrophobic drugs, could be tuned by grafting hydrophobic side chains or playing on the suspending medium parameters. Very importantly, new strategies for cross-linking, allowing one at the same time to preserve the biofunctionality of biomolecules, will have to be developed. This may include the development of photo-crosslinking, which is already widely used for polysaccharide hydrogels, or the preparation of composite films containing organic-inorganic molecules, that could in addition contain functionalized nanocolloids. The biological applications of LbL are still at an early stage but will probably see rapid development in the next decade. In particular, smart systems will be developed, making use of the ‘reservoir’ capacity of the polysaccharide multilayer films. These reservoirs will have to be optimized in order to increase the loaded amount, to optimize the release, while at the same time allowing for cell adhesion. Also, polysaccharide LbL will help to answer fundamental biology questions and to modulate cell adhesion, proliferation and differentiation, depending on the film properties. In this way, the films may also constitute a new surface for stem cells specification and/or differentiation. Surface transfection by means of LbL will be developed and optimized. Studies of the durability and tissue response in vivo will have to be conducted for films that are aimed at being implanted or in contact with human tissues. Finally, one can foresee that the films will not only be used on 2D flat surfaces, but also on 3D scaffolds. It can be expected that natural based multilayer films with embedded growth factors or specific ligands deposited on complex 3D scaffolds will help the development of tissue engineered matrices. In summary, I truly believe that the natural based films have a future in biological applications and will be used for the development of new biomaterials.

8.7

Sources of further information and advice

A website that contain all the references up to 2005: http://www.chem.fsu.edu/multilayers/

Key books 1. Multilayer Thin Films – Sequential Assembly of Nanocomposite Materials. Editors Gero Decher and Joseph B. Schlenoff, (Foreword by Jean-Marie Lehn) ISBN: 3-527-30440-1, Hardcover, 544 Pages, Wiley-VCH, February 2003.

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This book is the first book on multilayers and gives a very good introduction to the topic. 2. Macromolecular Engineering, Volume 2 ‘Elements of macromolecular structural control’ (chapter on polyelectrolyte multilayer films). Editors Matyjaszewski, K., Gnanou, Y., Leibler, L. Willey-VCH, ISBN; 978-2527-31446-1, 2007.

Reviews Ai H, Jones S A, Lvov Y M, (2003), Biomedical applications of electrostatic layer-by-layer nano-assembly of polymers, enzymes, and nanoparticles. Cell Biochemistry and Biophysics, 39, 23–43. Jaber J A, Schlenoff J B, (2006), Recent developments in the properties and applications of polyelectrolyte multilayers. Current Opinion in Colloid and Interface Science, 11, 324–329. Schonhoff M (2003), Self-assembled polyelectrolyte multilayers. Current Opinion in Colloid and Interface Science, 8, 86–95. Sukhishvili S A, Kharlampieva E, Izumrudov V, (2006), Where polyelectrolyte multilayers and polyelectrolyte complexes meet, Macromolecules, 39, 8873–8881. Sukhorukov G B, Mohwald H, (2007), Multifunctional cargo systems for biotechnology, Trends in Biotechnology, 25, 93–98. Tang Z Y, Wang Y, Podsiadlo P and Kotov N A, (2006), Biomedical applications of layer-by-layer assembly: From biomimetics to tissue engineering, Advanced Materials, 18, 3203–3224.

Scientific societies American Chemical Society (www.acs.org), Materials Research Society (www.mrs.org). Sessions on polyelectrolyte multilayer films are held at their annual meetings. The Gordon Research Conferences (GRC) also organizes a conference on ‘Organic Thin Films’ held every second year.

8.8

References

1 Whitesides G M and Boncheva M (2002), Beyond molecules: self-assembly of mesoscopic and macroscopic components, Proceeding of the National Academy of Science USA, 99, 4769–4774. 2 Hammond P T (1999), Recent explorations in electrostatic multilayer thin film assembly, Current Opinion in Colloid & Interface Science, 4, 430–442. 3 Bertrand P, Jonas A, Laschewsky A and Legras R (2000), Ultrathin polymer coatings by complexation of polyelectrolytes at interfaces : suitable materials, structure and properties, Macromolecular Rapid Communications, 21, 319–348.

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4 Decher G, Hong J D and Schmitt J (1992), Buildup of ultrathin multilayer films by a self-assembly process. Consecutively alternating adsorption of anionic and cationic polyeletrolytes on charges surface, Thin Solid Films, 1992: 831–835. 5 Decher G (1997), Fuzzy nanoassemblies: Toward layered polymeric multicomposites, Science, 277, 1232–1237. 6 Elbert D L, Herbert C B and Hubbell J A (1999), Thin polymer layers formed by polyelectrolyte multilayer techniques on biological surfaces, Langmuir, 15, 5355– 5362. 7 Serizawa T, Yamaguchi M and Akashi M (2002), Alternating bioactivity of polymeric layer-by-layer assemblies: Anticoagulation vs procoagulation of human blood, Biomacromolecules, 3, 724–731. 8 Shenoy D B, Antipov A, Sukhorukov G B and Möhwald H (2003), Layer-by-layer engineering of biocompatible, decomposable core-shell structures, Biomacromolecules, 4, 265–272. 9 Richert L, Lavalle P, Payan E, Stoltz J-F, Shu X Z, Prestwich G D, Schaaf P, Voegel J-C and Picart C (2004), Layer-by-layer buildup of polysaccharide films: Physical chemistry and cellular adhesion aspects, Langmuir, 1, 284–294. 10 Thierry B, Winnik F M, Merhi Y, Silver J and Tabrizian M (2003), Bioactive coatings of endovascular stents based on polyelectrolyte multilayers, Biomacromolecules, 4, 1564–1571. 11 Balabushevich N G, Tiourina O P, Volodkin D V, Larionova N I and Sukhorukov G B (2003), Loading the multilayer dextran sulfate/protamine microsized capsules with peroxidase, Biomacromolecules, 4, 1191–1197. 12 Cui W, Barr G, Faucher K M, Sun X L, Safley S A, Weber C J and Chaikof E L (2004), A membrane-mimetic barrier for islet encapsulation, Transplantation Proceedings, 36, 1206–1208. 13 Hubbell J A (1999), Bioactive biomaterials, Current Opinion in Biotechnology, 10, 123–129. 14 Suh J K F and Matthew H W T (2000), Application of chitosan-based polysaccharide biomaterials in cartilage tissue engineering: a review, Biomaterials, 21, 2589– 2598. 15 Ueno H, Mori T and Fujinaga T (2001), Topical formulations and wound healing applications of chitosan, Advanced Drug Delivery Reviews, 52, 105–115. 16 Laurent T C (1998), The Chemistry, Biology, and Medical Applications of Hyaluronan and its Derivatives. Cambridge, UK: Cambridge University Press. 17 Kumar M N V R (2000), A review of chitin and chitosan applications, Reactive and Functional Polymers, 46, 1–27. 18 Serizawa T, Yamaguchi M and Akashi M (2002), Enzymatic hydrolysis of a layerby-layer assembly prepared from chitosan and dextran sulfate, Macromolecules, 35, 8656–8658. 19 Schenkels L C, Veerman E C, Nieu W and Amerongen A V (1995), Biochemical composition of human saliva in relation to other mucosal fluids, Critical Review in Oral and Biological Medicine, 6, 161–175. 20 Menzel E J and Farr C (1998), Hyaluronidase and its substrate hyaluronan: biochemistry, biological activities and therapeutic uses, Cancer Letters, 131, 3–11. 21 Kirker K R, Luo Y, Nielson J H, Shelby J and Prestwich G D (2002), Glycosaminoglycan hydrogel films as bio-interactive dressings for wound healing, Biomaterials, 23, 3661–3671.

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22 Nettles D L, Elder S H and Gilbert J A (2002), Potential use of chitosan as a cell scaffold material for cartilage tissue engineering, Tissue Eng, 8, 1009–1016. 23 Ishihara M, Obara K, Ishizuka T, Fujita M, Sato M, Masuoka K, Saito Y, Yura H, Matsui T, Hattori H, Kikuchi M and Kurita A (2003), Controlled release of fibroblast growth factors and heparin from photocrosslinked chitosan hydrogels and subsequent effect on in vivo vascularization, Journal of Biomedical Materials Research, 64A, 551–559. 24 Fouissac E, Milas M, Rinaudo M and Borsali R (1992), Influence of the ionic strength on the dimensions of sodium hyaluronate, Macromolecules, 25, 5613– 5617. 25 Berth G and Dautzenberg H (2002), The degree of acetylation of chitosans and its effect on the chain conformation in aqueous solution, Carbohydrate Polymers, 47, 39–51. 26 Colfen H, Berth G and Dautzenberg H (2001), Hydrodynamic studies on chitosans in aqueous solution, Carbohydrate Polymers, 45, 373–383. 27 Desbrieres J, Martinez C and Rinaudo M (1996), Hydrophobic derivatives of chitosan: characterization and rheological behaviour, Int J Biol Macromol, 19, 21–28. 28 Sabnis S and Block L H (2000), Chitosan as an enabling excipient for drug delivery systems. I. Molecular modifications, Int J Biol Macromol, 27, 181–186. 29 Prestwich G D, Marecak D M, Marecek J F, Vercruysse K P and Ziebell M R (1998), Controlled chemical modification of hyaluronic acid: synthesis, applications, and biodegradation of hydrazide derivatives, J Control Release, 53, 93–103. 30 Luo Y and Prestwich G D (1999), Synthesis and selective cytotoxicity of a hyaluronic acid-antitumor bioconjugate, Bioconjugate Chemistry, 10, 755–763. 31 Morimoto M, Saimoto H, Usui H, Okamoto Y, Minami S and Shigemasa Y (2001), Biological activities of carbohydrate-branched chitosan derivatives, Biomacromolecules, 2, 1133–1136. 32 Machens H G, Berger A C and Mailaender P (2000), Bioartificial skin, Cells Tissues Organs, 167, 88–94. 33 Kosaraju L (2005), Colon targeted delivery systems: review of polysaccharides for encapsulation and delivery, Crit Rev Food Sci Nutr, 45, 251–258. 34 Ladam G, Schaad P, Voegel J-C, Schaaf P, Decher G and Cuisinier F J G (2000), In situ determination of the structural properties of initially deposited polyelectrolyte multilayers, Langmuir, 16, 1249–1255. 35 Ruths J, Essler F, Decher G and Riegler H (2000), Polyelectrolytes I: Polyanion/ polycation multilayers at the air/monolayer/water interface as elements for quantitative polymer adsorption studies and preparation of hetero-superlattices on solid surfaces, Langmuir, 16, 8871–8878. 36 Caruso F, Lichtenfeld H, Donath E and Mohwald H (1999), Investigation of electrostatic interactions in polyelectrolyte multilayer films: Binding of anionic fluorescent probes to layers assembled onto colloids, Macromolecules, 32, 2317– 2328. 37 Caruso F, Furlong D N, Ariga K, Ichinose I and Kunitake T (1998), Characterization of polyelectrolyte-protein multilayer films by atomic force microscopy, scanning electron microscopy, and Fourier transform infrared reflection-absorption spectroscopy, Langmuir, 14, 4559–4565. 38 Caruso F, Niikura K, Furlong D N and Okahata Y (1997), 1. Ultrathin multilayer polyelectrolyte films on gold: Construction and thickness determination, Langmuir, 13, 3422–3426.

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39 Picart C, Ladam G, Senger B, Voegel J-C, Schaaf P, Cuisinier F J G and Gergely C (2001), Determination of structural parameters characterizing thin films by optical methods: A comparison between scanning angle reflectometry and optical waveguide lightmode spectroscopy, Journal of Chemical Physics, 115, 1086–1094. 40 Picart C, Lavalle P, Hubert P, Cuisinier F J G, Decher G, Schaaf P and Voegel JC (2001), Buildup mechanism for poly(L-lysine)/hyaluronic acid films onto a solid surface, Langmuir, 17, 7414–7424. 41 Picart C, Mutterer J, Richert L, Luo Y, Prestwich G D, Schaaf P, Voegel J-C and Lavalle P (2002), Molecular basis for the explanation of the exponential growth of polyelectrolyte multilayers, Proceedings of the National Academy of Sciences of the United States of America, 99, 12531–12535. 42 Tezcaner A, Hicks D, Boulmedais F, Sahel J, Schaaf P, Voegel J C and Lavalle P (2006), Polyelectrolyte multilayer films as substrates for photoreceptor cells, Biomacromolecules, 7, 86–94. 43 Lavalle P, Gergely C, Cuisinie F, Decher G, Schaaf P, Voegel J-C and Picart C (2002), Comparison of the structure of polyelectrolyte multilayer films exhibiting a linear and an exponential growth regime: An in situ atomic force microscopy study, Macromolecules, 35, 4458–4465. 44 DeLongchamp D M, Kastantin M and Hammond P T (2003), High-contrast electrochromism from layer-by-layer polymer films, Chemistry of Materials, 15, 1575–1586. 45 McAloney R A, Sinyor M, Dudnik V and Goh M C (2001), Atomic force microscopy studies of salt effects on polyelectrolyte multilayer film morphology, Langmuir, 17, 6655–6663. 46 Boulmedais F, Ball V, Schwinté P, Frisch B, Schaaf P and Voegel J-C (2002), Buildup of exponentially growing multilayer polypeptide films with internal secondary structure, Langmuir, 19, 440–445. 47 Zhang J, Senger B, Vautier D, Picart C, Schaaf P, Voegel J-C and Lavalle P (2005), Buildup of collagen and hyaluronic acid polyelectrolyte multilayers, Biomaterials, 26, 3353–3361. 48 Johansson J A, Halthur T, Herranen M, Soderberg L, Elofsson U and Hilborn J (2005), Build-up of collagen and hyaluronic acid polyelectrolyte multilayers, Biomacromolecules, 6, 1353–1359. 49 Burke S E and Barrett C J (2003), pH-responsive properties of multilayered poly(Llysine)/hyaluronic acid surfaces, Biomacromolecules, 4, 1773–1783. 50 Ai H, Lvov Y, Mills D, Jennings M, Alexander J and Jones S (2003), Coating and selective deposition of nanofilm on silicone rubber for cell adhesion and growth, Cell Biochemistry and Biophysics, 38, 103–114. 51 Tryoen-Toth P, Vautier D, Haikel Y, Voegel J-C, Schaaf P, Chluba J and Ogier J (2002), Viability, adhesion, and bone phenotype of osteoblast-like cells on polyelectrolyte multilayer films, Journal of Biomedical Materials Research, 60, 657–667. 52 Lapcik L, Lapcik L, De Smedt S, Demeester J and Chabrecek P (1998), Hyaluronan: Preparation, structure, properties, and applications, Chemical Reviews, 98, 2663– 2684. 53 Richert L, Boulmedais F, Lavalle P, Mutterer J, Ferreux E, Decher G, Schaaf P, Voegel J-C and Picart C (2004), Improvement of stability and cell adhesion properties of polyelectrolyte multilayer films by chemical cross-linking, Biomacromolecules, 5, 284–294.

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54 Boulmedais F, Schwinté P, Gergely C, Voegel J C and Schaaf P (2002), Secondary structure of polypeptide multilayer films: An example of locally ordered polyelectrolyte multilayers, Langmuir, 18, 4523–4525. 55 Shiratori S S and Rubner M F (2000), pH-dependent thickness behavior of sequentially adsorbed layers of weak polyelectrolytes, Macromolecules, 33, 4213–4219. 56 Turner R, Lin P and Cowman M (1988), Self-association of hyaluronate segments in aqueous NaCl solution, Archives of Biochemistry and Biophysics, 265, 484–495. 57 Ghosh S, Kobal I, Zanette D and Reed W (1993). Conformational contraction and hydrolysis of hyaluronanate in sodium-hydroxyde solutions, Macromolecules, 26, 4685–4693. 58 Burke S E and Barrett C J (2005), Swelling behavior of hyaluronic acid/polyallylamine hydrochloride multilayer films, Biomacromolecules, 6, 1419–1428. 59 Yasui S and T Keigerling (1986), Vibrational circular-dichroism of polypeptides. Poly(lysine) conformations as a function of pH in aqueous-solution, Journal of the American Chemical Society, 108, 5576–5581. 60 Picart C, Schneider A, Etienne O, Mutterer J, Egles C, Jessel N and Voegel J-C (2005), Controlled degradability of polysaccharide multilayer films in vitro and in vivo, Advanced Functional Materials, 15, 1771–1780. 61 Rhazi M, Desbrieres J, Tolaimate A, Rinaudo M, Vottero P, Alagui A and El Meray M (2002), Influence of the nature of the metal ions on the complexation with chitosan. Application to the treatment of liquid waste, European Polymer Journal, 38, 1523–1530. 62 Amsden B and Turner N (1999), Diffusion characteristics of calcium alginate gels, Biotechnology and Bioengineering, 65, 605–610. 63 Kharlampieva E and Sukhishvili S A (2003), Ionization and pH stability of multilayers formed by self-assembly of weak polyelectrolytes, Langmuir, 19, 1235–1243. 64 Richert L, Arntz Y, Schaaf P, Voegel J-C and Picart C (2004), pH dependent growth of poly(L-lysine/poly(L-glutamic)) acid multilayer films and their cell adhesion properties, Surface Science, 570, 13–29. 65 Mao Z, Ma L, Zhou J, Gao C and Shen J (2005), Bioactive thin film of acidic fibroblast growth factor fabricated by layer-by-layer assembly, Bioconjugate Chemistry, 16, 1316–1322. 66 Ma L, Zhou J, Gao C and Shen J (2007), Incorporation of basic fibroblast growth factor by a layer-by-layer assembly technique to produce bioactive substrates, Journal of Biomedical Materials Research Part B, Applied Biomaterials, 83, 285– 292. 67 Discher D E, Janmey P and Wang Y L (2005), Tissue cells feel and respond to the stiffness of their substrate, Science, 310, 1139–1143. 68 Tan Q, Ji J, Barbosa M A, Fonseca C and Shen J (2003), Constructing thromboresistant surface on biomedical stainless steel via layer-by-layer deposition anticoagulant, Biomaterials, 24, 4699–4705. 69 Zhu Y, Gao C, He T, Liu X and Shen J (2003), Layer-by-Layer assembly to modify poly(L-lactic acid) surface toward improving its cytocompatibility to human endothelial cells, Biomacromolecules, 4, 446–452. 70 Vautier D, Hemmerlé Vodouhe C, Koenig G, Richert L, Picart C, Voegel J-C, Debry C, Chluba J and Ogier J (2003), 3d surface charges modulate protusive and contractile contacts of chondrosarcoma cells, Cell Motility and the Cytoskeleton, 56, 147–158. 71 Sukhorukov G B, Rogach A L, Garstka M, Springer S, Parak W J, Munoz-Javier

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A, Kreft O, Skirtach A G, Susha A S, Ramaye Y, Palankar R and Winterhalter M (2007), Multifunctionalized polymer microcapsules, novel tools for biological and pharmacological applications, Small, 3, 944–955. De Geest B G, Sanders N N, Sukhorukov G B, Demeester J and De Smedt S C (2007), Release mechanisms for polyelectrolyte capsules, Chemical Society Reviews, 36, 636–649. Donath E, Sukhorukov G, Caruso F, Davis S and Möhwald H (1998), Novel hollow polymer shells by colloid-templated assembly of polyelectrolytes, Angewandte Chemie International Edition, 37, 2201–2205. Caruso F, Yang W, Trau D and Renneberg R (2000), Microencapsulation of uncharged low molecular weight organic materials by polyelectrolyte multilayer self-assembly, Langmuir, 16, 8932–8936. Antipov A A, Sukhorukov G, Donath E and Möhwald H (2001), Sustained release properties of polyelectrolyte multilayer capsules, Journal of Physical Chemistry B, 105, 2281–2284. Qiu X, Leporatti S, Donath E and Möhwald H (2001), Studies on the drug release properties of polysaccharide multilayer encapsulated ibuprofen microparticles, Langmuir, 17, 5375–5380. Déjugnat C and Sukhorukov G (2004), pH-responsive properties of hollow polyelectrolyte microcapsules templated on various cores, Langmuir, 20, 7265– 7269. Tong W, Gao C and Möhwald H (2005), Manipulating the properties of polyelectrolyte microcapsules by glutaraldehyde cross-linking, Chemistry of Materials, 17, 4610– 4616. Berth G, Voigt A, Dautzenberg H, Donath E and Möhwald H (2002), Polyelectrolyte complexes and layer-by-layer capsules from chitosan/chitosan sulfate, Biomacromolecules, 3, 579–590. Zhang Y, Guan Y and Zhou S (2005), Single component chitosan hydrogel microcapsule from a layer-by-layer approach, Biomacromolecules, 6, 2365–2369. Itoh Y, Matsusaki M, Kida T and Akashi M (2006), Enzyme-responsive release of encapsulated proteins from biodegradable hollow capsules, Biomacromolecules, 7, 2715–2718. Kreft O, Prevot M, Mohwald H and Sukhorukov G B (2007), Shell-in-shell microcapsules: a novel tool for integrated, spatially confined enzymatic reactions, Angew Chem Int Ed Engl, 46, 5605–5608. Mamedov A A, Kotov N A, Prato M, Guldi D M, Wicksted J P and Hirsch A (2002), Molecular design of strong single-wall carbon nanotube/polyelectrolyte multilayer composites, Nature Materials, 1, 190–194. Jiang C and Tsukruk V (2006), Freestanding nanostructures via layer-by-layer assembly, Advanced Materials, 18, 829–840. Rmaile H H and Schlenoff J B (2003), Optically active polyelectrolyte multilayers as membranes for chiral separations, Journal of the American Chemical Society, 125, 6602–6603. Miller M D and Bruening M L (2005), Correlation of the swelling and permeability of polyelectrolyte multilayer films, Chemistry of Materials, 17, 5375–5381. Lavalle P, Boulmedais F, Ball V, Mutterer J, Schaaf P and Voegel J (2005), Free standing membranes made of biocompatible polyelectrolytes using the layer by layer method, Journal of Membrane Science, 253, 49–56.

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88 Podsiadlo P, Tang Z, Shim B S and Kotov N A (2007), Counterintuitive effect of molecular strength and role of molecular rigidity on mechanical properties of layer-by-layer assembled nanocomposites, Nano Letters, 7, 1224–1231. 89 Lu H and Hu NF (2006), Loading behavior of {chitosan/hyaluronic acid}(n) layerby-layer assembly films toward myoglobin: An electrochemical study, Journal of Physical Chemistry B, 110, 23710–23718. 90 Vodouhe C, Guen E L, Garza J M, Francius G, Dejugnat C, Ogier J, Schaaf P, Voegel J C and Lavalle P (2006), Control of drug accessibility on functional polyelectrolyte multilayer films, Biomaterials, 27, 4149–4156. 91 Schneider A, Picart C, Senger B, Schaaf P, Voegel J-C and Frisch B (2007), Layerby-layer films from hyaluronan and amine-modified hyaluronan, Langmuir, 23, 2655–2662. 92 Schneider A, Vodouhé A, Richert L, Francius G, Le Guen E, Schaaf P, Voegel J-C, Frisch F and Picart C (2007), Multi-functional polyelectrolyte multilayer films: combining mechanical resistance, biodegradability and bioactivity, Biomacromolecules, 8, 139–145. 93 Dimitrova M, Arntz Y, Lavalle P, Meyer F, Wolf M, Schuster C and Haikel Y (2007), Adenoviral gene delivery from multilayered polyelectrolyte architectures, Advanced Functional Materials, 17, 233–245. 94 Thierry B, Kujawa P, Tkaczyk C, Winnik F M, Bilodeau L and Tabrizian M (2005), Delivery platform for hydrophobic drugs: prodrug approach combined with selfassembled multilayers, Journal of the American Chemical Society, 127, 1626– 1627. 95 Croll T I, O’Connor A J, Stevens G W and Cooper-White J J (2006), A blank slate? Layer-by-layer deposition of hyaluronic acid and chitosan onto various surfaces, Biomacromolecules, 7, 1610–1622. 96 Fu J, Ji J, Yuan W and Shen J (2005), Construction of anti-adhesive and antibacterial multilayer films via layer-by-layer assembly of heparin and chitosan, Biomaterials, 26, 6684–6692. 97 Etienne O, Schneider A, Taddei C, Richert L, Schaaf P, Voegel J-C, Egles C and Picart C (2005), Degradability of polysaccharide multilayer films in the oral environment: an in vitro and in vivo study, Biomacromolecules, 6, 726–733. 98 Richert L, Schneider A, Vautier D, Jessel N, Payan E, Schaaf P, Voegel J-C and Picart C (2006), Imaging cell interactions with native and cross-linked polyelectrolyte multilayers, Cell Biochemistry and Biophysics, 44, 273–276. 99 Etienne O, Picart C, Taddei C, Haikel Y, Dimarcq J-L, Schaaf F, Voegel J-C, Ogier J A and Egles C (2004), Multilayer polyelectrolyte films functionalized by insertion of defensin: a new approach to protection of implants from bacterial colonization, Antimicrobial Agents and Chemotherapy, 48, 3662–3669. 100 Schoeler B, Delorme N, Doench I, Sukhorukov G B, Fery A and Glinel K (2006), Polyelectrolyte films based on polysaccharides of different conformations: effects on multilayer structure and mechanical properties, Biomacromolecules, 7, 2065– 2071. 101 Schneider A, Bolcato-Bellemin A-L, Francius G, Jedrzejwska J, Schaaf P, Voegel J-C, Frisch B and Picart C (2006), Glycated polyelectrolyte multilayer films : differential adhesion of primary versus tumor cells, Biomacromolecules, 8, 139– 145. 102 Sinani V A, Koktysh D S, Yun B-G, Matts R L, Pappas T C, Motamedi M, Thomas S N and Kotov N A (2003), Collagen coating promotes biocompatibility of

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semiconductor nanoparticles in stratified LBL films, Nano Letters, 3, 1177–1182. 103 Koktysh D S, Liang X, Yun B G, Pastoriza-Santos I, Matts R L, Giersig M, SerraRodrîguez C, Liz-Marzán L M and Kotov N A (2002), Biomaterials by design: Layer-by-layer assembled ion-selective and biocompatible films of TiO2 nanoshells for neurochemical monitoring, Advanced Functional Materials, 12, 255–265. 104 Francius G, Hemmerle J, Voegel J C, Schaaf P, Senger B and Ball V (2007), Anomalous thickness evolution of multilayer films made from poly-L-lysine and mixtures of hyaluronic acid and polystyrene sulfonate, Langmuir, 23, 2602–2607. 105 Richert L, Engler A J, Discher D E and Picart C (2004), Elasticity of native and cross-linked polyelectrolyte multilayers, Biomacromolecules, 5, 1908–1916. 106 Silverstein R M and Webster F X (1997), Infrared spectrometry. In: Silverstein R M, Webster F X and Kiemle D J (eds) Spectrometric Identification of Organic Compounds. 6th edition New York: Wiley, 72–126. 107 Tomihata K and Ikada Y (1997), Crosslinking of hyaluronic acid with watersoluble carbodiimide, J Biomed Mater Res, 37, 243–51. 108 Francius G, Hemmerle J, Ohayon J, Schaaf P, Voegel J-C, Picart C and Senger B (2006), Effect of cross-linking on the elasticity of polyelectrolyte multilayer films measured by colloidal probe AFM, Microscopy Research and Techniques, 69, 84– 92. 109 Schneider A, Richert L, Francius G, Voegel J-C and Picart C (2007), Elasticity, biodegradability and cell adhesion properties of chitosan/hyaluronan multilayer films, Biomedical Materials and Engineering, 2, 1–7. 110 Picart C, Elkaim R, Richert L, Audoin F, Da Silva Cardoso M, Schaaf P, Voegel JC and Frisch B (2005), Primary cell adhesion on RGD functionalized and covalently cross-linked polyelectrolyte multilayer thin films, Advanced Functional Materials, 15, 83–94. 111 Schneider A, Francius G, Obeid R, Schwinté P, Frisch B, Schaaf P, Voegel J-C, Senger B and Picart C (2006), Polyelectrolyte multilayer with tunable Young’s modulus : influence on cell adhesion, Langmuir, 7, 2882–2889. 112 Boulmedais F, Tang C S, Keller B and Voros J (2006), Controlled electrodissolution of polyelectrolyte multilayers: A platform technology towards the surface-initiated delivery of drugs, Advanced Functional Materials, 16, 63–70. 113 Kujawa P, Moraille P, Sanchez J, Badia A and Winnik F M (2005), Effect of molecular weight on the exponential growth and morphology of hyaluronan/chitosan multilayers: a surface plasmon resonance spectroscopy and atomic force microscopy investigation, Journal of the American Chemical Society, 127, 9224–9234. 114 Svensson O, Lindh L, Cardenas M and Arnebrant T (2006), Layer-by-layer assembly of mucin and chitosan–Influence of surface properties, concentration and type of mucin, Journal of Colloid and Interface Science, 299, 608–616.

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9 Peptide modification of polysaccharide scaffolds for targeted cell signaling S. L É V E S Q U E, R. W Y L I E, Y. A I Z A WA and M. S H O I C H E T, University of Toronto, Canada

9.1

Introduction

Polysaccharides are excellent substrates as biomimetic scaffolds for use in tissue engineering and regenerative medicine since most are biologically inert and non-cell-adhesive, allowing these functions to be controlled through modification. Through specific modification with cell-signaling factors, such as peptides, proteins and growth factors, the cellular microenvironment can be defined. By engineering this environment – chemically, physically and mechanically – cell behavior can be controlled and tuned to specific functions, thereby leading to tissue organization and function. This chapter reviews the rationale for controlling the cellular microenvironment and the polysaccharides studied therein for ultimate use in tissue engineering and regeneration.

9.1.1

Impact of the extracellular environment on cell signaling

Cell signaling can be defined as the exchange of information or signals between the cell and its extracellular environment. This dialogue results in a wide variety of cell-specific signaling pathways, known as signal transduction, that ultimately regulates a number of complex biological processes such as cell differentiation, proliferation, migration, expression of other genes, or apoptosis.1 In mammalian tissues, the extracellular microenvironment is a highly hydrated network made up of three main components: (1) the ECM, comprised of a complex 3-D scaffold of structural proteins such as collagens and elastin; (2) soluble cell-secreted macromolecular signals including growth factors, chemokines and cytokines; and (3) adhesive glycoproteins such as laminin, fibronectin, vitronectin, tenascin, and hydrophilic proteoglycans. Surface proteins of neighboring cells consist of integrins and other cell adhesive molecules (CAMs).1 These three main components provide different signals, which regulate cell behavior. 260 © 2008, Woodhead Publishing Limited

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Growth factors, chemokines and cytokines are cell-secreted proteins capable of stimulating cell proliferation, differentiation and other cellular behaviors. They play a crucial role in cell signaling by binding to specific cell surface receptors, triggering transduction. The ECM acts as a reservoir of these soluble signals, where they can be sequestered within, bound to its proteoglycans.2,3 By harboring growth factors and cytokines, the ECM influences cell behavior in many ways. The direct binding of growth factors to the matrix can affect their local concentration,4 diffusion,5 concentration gradient6 or general availability to cell receptors. Moreover, biological activity is preserved by either protecting against protease degradation7 or by presenting them to the cell-surface receptors.3,5,8 Thus the ECM provides a microenvironment of biologically active molecules that regulate cell signaling. Cells in tissues bind to each other (cell-cell adhesion) and anchor to the ECM (cell-matrix adhesion) through the binding of CAMs. Cell adhesion mechanisms not only provide structural integrity to tissue but also play an active role in events influencing cell fate by triggering different cellular signaling pathways. The majority of CAMs are regrouped into four distinct families: cadherins, immunoglobulin superfamily, integrins and selectins.9 The integrin family has cell-surface receptors primarily involved in cellmatrix adhesion, binding to specific amino-acid motifs found in ECM proteins. One of the main roles of integrins is cellular adhesion to the ECM, linking the cell cytoskeleton to the ECM fibers. The integrins bind to ECM, forming clusters of focal adhesion and promote the assembly of actin filaments via association with cytoskeletal and signaling complexes. The actin filaments reorganize themselves into larger stress fibers, causing more integrin clustering, thereby enhancing the matrix binding and organization by integrins in a positive feedback system.10 Integrins are not merely mediators for cell adhesion. The reorganization of the cytoskeletal network, following integrin binding to the ECM, activates signaling molecules which in turn bind to the network and attract additional signaling molecules, forming a complex feedback loop. Integrins are also involved in crosstalk with other receptors.10,11 The integrin-cytoskeletal complexes play a critical role linking growth-factor receptors and G-protein coupled receptors (GPCRs) to signaling proteins. For example, the integrin engagement with the ECM allows the activation of growth-factor receptors such as epidermal growth factor receptor (EGF-R), platelet-derived growth factor receptor (PDGF-R), and vascular endothelial growth factor receptor (VEGF-R) while the detachment from the ECM inhibits the growth-factor activation of focal adhesion kinase (FAK).11 Integrin signals are required for the growth-factor activation of extracellular signal-regulated kinase (ERK), influencing cell cycle progression.11 In addition, integrins have been found

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to induce receptor expression. In several cell types, growth-factor receptor enhancement of motility resulted from the upregulation of integrin receptor expression.11 In addition to the chemical stimuli provided by soluble cytokines and ECM proteins, mechanical forces also induce changes in cell behavior. External forces, such as tensile, compressive, or shear forces, applied to the extracellular matrix and cells have been reported to regulate cell functions, including gene induction, protein synthesis, cell growth and differentiation. Integrins play an important role in this process called mechanotransduction. They not only serve as adhesive receptors 12 but also as mechanoreceptors or mechanosensors.10,13–15 Being transmembrane proteins, they physically link the cytoskeleton to the ECM proteins, relaying external physical stimuli to the cell cytoskeleton. A wide variety of effects have been observed in multiple types of cells in response to mechanical signals. For example, mechanical forces have been reported to promote, in chondrocytes, the secretion of autocrine and paracrine acting soluble mediators,16 and gene expression of ECM components such as aggrecan17 and type II collagen.18 Mechanical stimuli have an influence on gene expression19 proliferation and differentiation of osteoblast cells.20,21 In addition, externally applied loads can lead to the expression of distinct ECM components, such as Types I and III collagen, fibronectin22–24 and an increase in fibroblast proliferation.22 Shear stress can change cell morphology and modulate signal transduction pathways and gene expression in vascular endothelial cells,25 induce directional migration26,27 and expression of matrix metalloproteinases.28 Smooth muscle cells have been reported to influence gene expression and synthesis of ECM components29,30 and matrix metalloproteinases.31 In the past, artificial biomaterial scaffolds were designed to support cell and tissue growth to match the macroscopic properties of the tissues and organs to be replaced without considering and recreating, on a nanoscale level, the cellular microenvironment observed in native tissues. In recent years, advances in molecular biology have provided a better understanding of the cellular environment which has been used to design systems that mimic the physical and biochemical interactions of the ECM with the cell. In this complex milieu, the ECM is clearly not just a simple support; it is also a critical component of all tissues providing a substructure for cell adhesion and movement and a storage depot for growth factors, chemokines, and cytokines, thereby playing a key role in cell signaling. Because the ECM influences cell fate, building scaffolds that can reproduce different aspects and functions of the ECM is a major focus in tissue engineering. Considerable effort has been focused on the development of multi-component systems that can elicit specific biological functions to spatially and temporally guide tissue regeneration. One way to mimic the ECM is to create artificial matrices modified with bioactive, cell-signaling peptides

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that comprise specific motifs, found in one or many glycoproteins, which are known to recognize and bind to cell-surface receptors such as integrins and to promote different behaviors depending on the cell type. If these peptides bind to integrins, the latter can act as mechanoreceptors and physical stimuli can be used to induce mechanotransduction. The immobilized peptides can then either mimic regions of growth factors or other chemokines or play a role in the binding and delivery of soluble growth factors.

9.1.2

Rationale of polysaccharide scaffolds

The architecture of the natural ECM has inspired several researchers to produce scaffolds that mimic several physical and/or biomechanical aspects of the ECM. One function of the scaffold is to provide an environment in which cells can infiltrate, proliferate, and express genes in order to form new tissue. The design of scaffold materials is governed by three principles: physical properties, mass transport and biological interactions.32 These properties or design variables are specific to the intended application and will often dictate the type of material used. Both natural and synthetic polymers have been used for tissue-engineering applications. The physical properties of the scaffolds include mechanical strength, morphology and degradation behavior. Many scaffolds initially fill a space otherwise occupied by natural tissue, in turn providing support for tissue regeneration. Macroscopically, the scaffold should provide temporary support sufficient to withstand in vivo forces and maintain a potential space for tissue development. In addition it must have mechanical properties similar to that of the targeted tissue. For example, the Young’s modulus of a scaffold for cortical bone would be between 3 and 30 GPa,33 for cartilage, between 0.7015.3 MPa,33 and for spinal cord, between 0.2-0.6 MPa.34,35 A mechanical mismatch between the surrounding tissue and the scaffold may damage the surrounding tissue and compromise the efficacy of the tissue engineering strategy. On the microscopic level, evidence suggests that the scaffold acts similarly to natural ECM. It transmits mechanical input to the cells which triggers mechanotransduction leading to cell growth, differentiation, and ultimately tissue formation.36–38 Scaffolds are usually only temporary matrices and either degrade via hydrolysis or enzymatic action, or dissolve, ideally after the new tissue has been generated. Many polysaccharide scaffolds, such as alginate, methyl cellulose, hyaluronan will dissolve under physiological conditions or be degraded by ectoenzymes such as lysozyme, hyaluronidase, or dextranase. Advantageously, many polysaccharide scaffolds (especially those derived from plants) can have the degradation rate and mechanism tailored by incorporation of degradable linkers which are either hydrolytically or enzymatically labile.39 Furthermore, polysaccharide scaffolds are considered

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biocompatible because they are both cytocompatible and do not elicit a chronic inflammatory response.32 For the creation of 3D tissues, it is important that the scaffolds allow and promote cell penetration and appropriate mass transport of nutrients into, and waste products out of, the matrix.32 This property is closely associated with scaffold morphology, dimensions (thickness) and vascularity. Macroporous scaffolds are highly desirable for the regeneration of specific tissues where the porosity and pore size distribution of the supporting 3-D structure impacts tissue formation.40 The porosity must also be continuous within the matrix to promote cell communication and effective tissue building.41 A number of methods have been examined to promote porosity, including phase separation,42,43 high pressure gas infiltration,41 and salt leaching42 among others.

9.1.3

Rationale of peptide modification

Scaffolds define the cellular microenvironment mechanically, physically and chemically. To guide cell response, scaffolds have been modified with ECMderived proteins and peptides.8 Cell-surface receptors, such as integrins, bind to different components of the ECM and some of the active peptide sequences of these ECM proteins have been identified and associated with specific cellular responses. Some of these peptide motifs are unique to one glycoprotein whereas others are more ubiquitous. The most studied amino acid sequence is the RGD motif, which was identified in fibronectin by Ruoslahti and Pierschbacher, as the first minimal essential peptide sequence to promote cell adhesion.44–46 Since then, the RGD sequence has also been identified in many other glycoproteins including fibronectin, collagens, laminins, tenascin, thrombospondin and vitronectin.6,47 Understanding that the ECM has an important role in cell behavior, many components of the ECM have been scrutinized to identify domains that interact with integrins or other cell-surface receptors. Numerous sequences have been identified that bind to cell-surface receptors and promote specific cellular responses. There are advantages and disadvantages of linking just the active peptide sequence to a polymeric scaffold vs. the entire ECM protein. While ECM proteins are capable of stimulating cells in multiple ways, defining precise cellular responses is better achieved with specific peptide sequences. Moreover, by immobilizing specific peptides, their conformation and concentration can be well-controlled which is not always easy with full proteins which can denature when immobilized and may not expose their active regions for facile cell binding. Peptides are commonly used instead of proteins since they are less susceptible to enzyme degradation, do not denature and are less sensitive to variations in pH and temperature.48

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265

Polysaccharide scaffolds in tissue engineering

The ideal scaffold mimics the extracellular matrix, defining an environment for cells that has the appropriate properties to support cell viability, function, tissue formation and host tissue integration. Biomaterials (and most often polymeric biomaterials) have been designed to serve as temporary replacements of the ECM for tissue regeneration in vitro and in vivo. Both naturallyderived and synthetic polymeric materials have been investigated as scaffolds, with great emphasis placed on the morphology, porosity (and pore size distribution), mechanical and chemical properties. Polysaccharides derived from natural sources, such as chitosan, alginate, hyaluronan, cellulose, agarose and dextran, are gaining appeal because they are largely biocompatible and biodegradable, often by enzymatic activity. Hydrogels are commonly used in tissue engineering because of their high water content, their usual non-cell adhesive nature and their ease of fabrication. Additionally, polysaccharide hydrogels are appealing as scaffolds because of the presence of functional groups for the attachment of signaling factors such as peptides. The properties of the polysaccharide hydrogels can be tailored by incorporation of chemical cross-linking agents and peptides. There are numerous examples of polysaccharides in tissue engineering. For example, hyaluronan has been used clinically in cartilage repair for many years49–53 while alginate and chitosan have been tested in bone,54 cartilage,43 and nerve55 regeneration strategies. Starch-based polymers and their derivatives have also been investigated as potential biomedical materials, especially in the orthopedic fields.56 Many of the polysaccharides used in tissue engineering are summarized in Table 9.1. In order to match the morphology of the scaffold to that of the targeted tissue, various processing methods have been developed with the goal of controlling both the properties on the macroscopic scale – i.e. overall shape, modulus, density, porosity – and microscopic scale – i.e. pore size, distribution and interconnectivity.72 Porosity is particularly important in scaffold design for facile penetration of cells and nutrients to accelerate regeneration in vivo. Moreover, a highly porous microstructure with interconnected porous networks is critical for cell-cell communication which is critical for cell survival, proliferation and migration. Several fabrication techniques have been developed for scaffold production, including: fiber bonding,66 phase separation,42,43 solvent casting,73 particulate leaching,42 membrane lamination,74 melt molding,75 gas foaming76/high pressure processing,41 freeze-drying77 and combinations of these techniques (e.g., gas foaming/particulate leaching, etc). For example, Zmora et al.,78 used freeze-drying to create the porous architecture in 3-D alginate scaffolds. They showed that the pore size affected both the compressibility of the

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Table 9.1 Commonly used polysaccharides in tissue engineering Provenances Agarose OH OH

O

O

H

O

O OH

O

O

Hyaluronan H

O

O

COOH O

O

O H n

NHCOCH3 OH OH OH

References

Spinal cord

57–59

Cartilage Skin Neural tissues Adipose tissues

49–53, 60–64

Skin Myocardial and cardiac tissues Liver Bone and cartilage Neural tissues

43, 65–76

Bone

56, 65–67

Neural tissues Vascular tissues

39, 68–71

Neural tissues Bone and cartilage Liver Ligaments and tendons Skin Vascular tissues

43, 54, 55, 72, 84–96

H n

OH

OH OH

Targeted tissue

Alginate

HOOC OH OH O OH O HO OH M

OH G HO HOOC M–M–M; M–G–M; G–G–G Starch OH O

HO

O

O OH HO

OH

O HO

O n

Dextran O HO HO

O H

O

H

HO

OH

O

OH OH O

Chitosan OH O O HO

O NH2 n

materials and hepatocyte morphogenesis. Gomes et al.66 used fiber bonding to fabricate starch-based fiber scaffolds and seeded marrow stromal cells on scaffolds of varying porosity: higher porosity enhanced cell proliferation and differentiation toward the development of bone-like mineralized tissue. Both freeze-drying and fiber bonding methods promote inter-connectivity of pores within the scaffold.66,78 While the above-mentioned techniques have shown significant results in academic laboratories, the techniques are limited by scale-up and

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reproducibility.72 Moreover, many of these scaffolds are difficult to sterilize or require potentially toxic organic solvents during their processing. Incomplete removal of toxic solvents from the fabricated scaffolds especially thicker constructs will result in harmfully residues that have adverse effects on adherent cells, incorporated biologically active agents or nearby tissues.72,79 Consequently, computer-controlled fabrication techniques, such as solid free-form fabrication (SFF) or rapid prototyping (RP), have recently gained popularity. Scaffolds synthesized by these techniques are constructed layerby-layer, resulting in a defined three-dimensional construct. In RP, various techniques have been developed to produce porous scaffolds79 and include: fused deposition modeling (FDM), 3D printing (3-DP),80 robocasting,81 and 3D bioplotting.82 Although most techniques have used synthetic polymers, Landers et al. created agarose scaffolds with well-defined internal pore structures using the 3D bioplotter system.82 Building on this technique, Ang et al.83 demonstrated that chitosan scaffolds with pore sizes between 400 and 1000 µm could be synthesized using a rapid prototyping robotic dispensing system, which consists of a computer-guided desktop robot and a onecomponent pneumatic dispenser. In addition to precisely controlling the porosity of the scaffold, 3-DP can also be used to manipulate the scaffold’s mechanical strength through repetitive deposition and processing of different materials in the layers. With 3-DP, successive 2D profiles are printed on freshly laid layers of powder until the scaffold is completed.80 By using this 3-DP technique, Lam et al.80 enhanced the mechanical strength of their scaffold by blending starch-based powders (cornstarch, dextran and gelatin) into the cylindrical architecture of the scaffold. While the RP system has many advantages in terms of scaffold design, it is still limited to specific materials84 and given pore sizes due to the resolution of the machine tools. More recently light-activated polysaccharide hydrogel scaffolds have been created where the 3D design has been dialed in by photo-activated chemical groups, resulting in well-defined chemical scaffolds. Unlike other scaffolds, which provide a series of inter-connected pores (or void volumes), the photoactivated scaffolds surround infiltrating cells with stimuli in 3D.58,85,86 Furthermore, the hydrogels are mechanically weak enough to allow cellular mobility without the incorporation of pores.58,87 This technique has been shown to be highly reproducible in both agarose and hyaluronan hydrogels and is not limited to the z-stacking processing of layer-by-layer deposition techniques.

9.3

Peptide immobilization

Peptide-modification of polymeric scaffolds has been investigated as biomimetic ECM for specific cell signaling. The binding of integrins to ECM proteins is mediated by cell-surface receptors with specific motifs

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found on the glycoprotein backbone and encoded by specific amino acid sequence. Given that the ECM has an important role in controlling cell behavior, many components of the ECM have been scrutinized to identify domains that interact with integrins or other cell-surface receptors. Numerous sequences were identified to recognize and bind to cell-surface receptors and to promote different behaviors depending on cell type. Table 9.2 presents the sequence and function of some of the peptides that have been derived from ECM proteins.

9.3.1

Peptide immobilization: Covalent versus adsorption

Peptides can be immobilized onto polysaccharide scaffolds through physical adsorption or covalent bonds. Adsorption between the peptide and polysaccharide occurs due to weak physical interactions such as van der Waals, electrostatic, ionic and hydrophobic interactions. Adsorption usually results in short term immobilization where the peptide will be released from the scaffold over time depending on the mechanism of adsorption and the strength of interactions. Covalent immobilization offers greater stability and control because it creates a covalent bond between the polysaccharide polymer chain and the peptide, providing long-term attachment. The concentration of the immobilized peptide can be controlled by varying the immobilization chemistry, the number of reactive sites on the polysaccharide or the peptide concentration in solution. Peptide immobilization onto polysaccharides Functional groups on polysaccharides are used as attachment sites for peptides and therefore will determine the method of immobilization. Polysaccharides can be categorized based on their functional groups: hydroxyls (agarose and dextran), carboxylic acids (alginate and hyaluronan), and amines (chitosan). Immobilization methods will be discussed in the context of these functional groups. Immobilization onto hydroxyl groups Peptides can be covalently bound to hydroxyl-containing polysaccharides through one-step or multi-step methods depending on the selected approach. The most common method involves the activation of the polysaccharide with carbonyldiimidazole (CDI) or disuccinimidyl carbonate (DSC) for reaction with primary amines at the N-terminus of the peptide.109,110 CDI reacts with hydroxyl groups to yield the reactive imidazole carbamate intermediate. The primary amine of a peptide can then react with the intermediate producing a

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Table 9.2 Active peptide sequences derived from ECM proteins Motif

Protein

Function

References

RGD

Fibronectin, Vibronectin, Laminin

Cell adhesion

69, 88–92

KQAGDV

Fibrinogen

Neurite extension Smooth muscle cell 91, 93, 94 adhesion

YIGSR

Laminin

Cell adhesion

69, 92, 95, 96

REDV

Fibronectin

Endothelial cell adhesion

93, 97

IKVAV

Laminin

Cell adhesion Neurite extension

69, 92, 95, 96, 98

ATLQLQEGRLHFXFDLGKGR

Murine Laminin

Fibroblast adhesion 93

GEFYFDLRLKGDKY

Human Collagen type IV

Fibroblast adhesion

93

YAVTGRGDSPAS

Human Fibronectin

Fibroblast adhesion Neurite outgrowth

93

KPSR

Heparin binding domain

Osteoblast adhesion

88, 99

VFDNFVLK

Tenascin-C

Neurite outgrowth

100

DINPYGFTVSWMASE

Tenascin-C

Neurite outgrowth

101

KNNQKSEPLIGRKKT

Fibronectin

Cell adhesion and neurite outgrowth

102

KDI

Laminin-1

Neurite outgrowth

103

YFQRYLI

Laminin-1

Cell adhesion and neurite outgrowth

104

IKLLI

Laminin-1

Cell adhesion and neurite outgrowth

104

KNRLTIELEVRT

Laminin-1

Cell adhesion and neurite outgrowth

105

KNSFMALYLSKG

Laminin-5

Cell adhesion and neurite outgrowth

106

GNSTISIRAPVY

Laminin-5

Cell adhesion and neurite outgrowth

106

KHIFSDDSSE

Neural cell adhesion molecule

Astrocyte adhesion

107

AGTFALRGDNPQG

Laminin

Cell adhesion

108

RKRLQVQLSIRT

Laminin

Cell adhesion

108

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stable carbamate linkage between the peptide and polysaccharide. Careful consideration must be taken when the peptide contains two or more primary amines to alleviate side reactions and crosslinking. DSC will also react with hydroxyl groups to form the succinimidyl carbonate for reaction with primary amines. The other method for peptide attachment involves the activation of the peptide with an isocyanate group. Isocyanates will react with hydroxyls forming a carbamate linkage (Figure 9.1). This method is limited because of the possible side reactions. Isocyanates will react with both hydroxyls and amines. Therefore the peptide should not contain any free primary amines or hydroxyls. All of the reactions mentioned above are water sensitive and should be performed under anhydrous conditions. Immobilization onto carboxylic acidsm (Fig. 9.2) 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDC) is the most common crosslinker for carboxylic acids and amines in water.111 EDC, in the presence of a carboxylic acid, will form an active ester intermediate which can then react with a primary amine to form an amide linkage. The major advantage of this reaction is that it can be performed in an aqueous environment. However the peptide can self-polymerize in the presence of EDC if it contains both a carboxylic acid and a primary amine. This can be alleviated by first activating the carboxylic acid containing polysaccharide with EDC, then adding the peptide once the activation step is complete. Sulfo-NHS (N-hydroxy Succinimide) can also be added to the reaction to increase the stability of the active ester intermediate, O-acylisourea, and thus the efficiency of this reaction. Other carbodiimide crosslinkers such as DIC and DCC are similar to EDC but are typically only used in organic solvents because they are insoluble in water. Carbonyl diimidazole also activates carboxylic acids for reaction with amines or thiols producing carbamate or thioester linkages, respectively.112 This method is interesting since no additional atoms are incorporated into the final product. The carboxylate reacts with CDI producing a N-acylimidazole, which then reacts with an amine or thiol of the peptide. Diazoalkanes and diazoacetyl compounds will also react with carboxylic acid groups. In this method the peptide must be activated, resulting in a diazoacetate group incorporated into the peptide. When mixed, the diazoacetate group of the peptide will spontaneously react with the carboxylic acid group of the polysaccharide. Immobilization onto amines As mentioned above, EDC can be used to link amines to carboxylic acids. In this case the amine would be on the polysaccharide (i.e. chitosan) and the

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O N

N N Carbonyl diimidazole

O Polysaccharide

O

H2N–Peptide

O N

Polysaccharide–OH

O

O

O

O

H2N–Peptide

O

Disuccinimidyl carbonate Polysaccharide

O

O

N

C

N

O

Peptide Polysaccharide

O

N H

O N Peptide H Carbamate linkage

Polysaccharide O

O

O

N

O O

N Peptide H Carbamate linkage

Polysaccharide

N N

O

O

Peptide

Carbamate linkage

9.1 Peptide immobilization onto hydroxyl containing polysaccharides.

O

Peptide modification of polysaccharide scaffolds

N

271

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carboxylate on the peptide C-terminus. The peptide can also self-polymerize in these condition since it contains both amines and carboxylic acids. Crosslinkers such as CDI and DCC can also be used to link amines and carboxylic acids, although these molecules can also induce peptide selfpolymerization and crosslink the polysaccharide by forming urea bonds. Therefore the amine group is usually first modified to contain a thiol group, as explained below. Methods involving thiols Maleimide and sulfhydryl groups are known to selectively react together producing a covalent linkage. Therefore the polysaccharide can be modified to contain the sulfhydryl group and the peptide the maleimide, or vice versa. Sulfhydryl groups can be incorporated into hydroxyl and amine containing polysaccharides using 2-iminothiolane.113 The N-terminus of the peptide can be modified with 3-maleimidopropionic acid to produce a terminal maleimide group.58 A maleimide group cannot be added to peptides that contain unprotected cysteines because the cysteine-thiol group will react with the maleimide.

9.4

Measurement

There are several ways to determine the amount of peptide immobilized: fluorescence, amino acid analysis and radioactivity. Peptide concentration can be determined through fluorescent spectroscopy by incorporating a fluorescent tag such as coumarin, fluorescein or rhodamine into the peptide sequence.58 The amount of fluorescence can be directly correlated to the peptide concentration. Similarly the peptide can be labeled with a radioactive isotope, e.g. tyrosine can be labeled with radioactive I-125.114,115 Amine reactive reagents such as CBQCA are used to determine peptide concentration by measuring the amount of primary amines present.116 CBQCA will form a fluorescent complex with primary amines, which can be quantified to determine the concentration of amines. By knowing the number of primary amines in the peptide, it is possible to deduce the peptide concentration. Ellman’s reagent can be used similarly by determining the amount of cysteines present.117 The BCA protein assay kit (Pierce) can determine peptide concentration through a colorimetric test.118 Another quantification method, amino acid analysis, hydrolyzes the peptides to their amino acids and detects the concentration of each amino acid using HPLC.119,120 The peptide concentration can then be determined though a series of comparisons. For example, if the peptide contains one glycine then the concentration of the peptide is equal to the concentration of glycine.

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O

EDC

Polysaccharide

O

H2N–Peptide

N

Polysaccharide

O N

H N

Peptide

Amide bond

O

Polysaccharide–COOH N

N

N N Carbonyl diimidazole

O Polysaccharide

Polysaccharide

N N

–

N

Peptide Diazoacetate peptide

O Peptide Polysaccharide

O

H N

Amide bond

O N+

O

H2N–Peptide

N H

Ester bond

9.2 Peptide immobilization onto carboxylic acid containing polysaccharides.

Peptide

Peptide modification of polysaccharide scaffolds

NH+

273

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9.5

Challenges associated with peptide immobilization

9.5.1

Peptide activity

The activity of immobilized peptides is dependent on their ability to bind with their target cell-membrane receptors. Peptides can absorb to the surface of polysaccharide scaffold hindering their ability to interact with receptors. Furthermore the peptides can adopt a conformation that is inactive and cannot bind with the receptor. This can be alleviated by incorporating amino acids found adjacent to the active site into the peptide sequence. These extra amino acids will allow the peptide to adopt a conformation closer to that found in the native protein. For example, surfaces modified with extended laminin-derived oligopeptides, CDPGYIGSR and CQAASIKVAV, which better maintain the native region conformation of the active sites, promoted a greater cellular response than those modified with the respective minimal peptide sequences.95 PEG and repetitive sequences of non-specific peptides, such as GGGGGG, have been used as spacers between the peptide sequence and the material surface.121,122 Another method to increase activity is to have multiple binding sites per peptide chain. For instance, GRGDSGRGDSGRGDSGRGDS has four possible binding sites, increasing the probability that a receptor will interact with the binding site of the peptide.123

9.5.2

Peptide concentration

Peptide concentration is also an important consideration. It is well accepted that only a small number of ECM-derived peptides is sufficient to promote cell adhesion. However Massia and Hubbell showed that the peptide concentration present at the material surface can influence cellular functions; a RGD density of 0.1 fmol/cm2 induced cell adhesion of fibroblasts but densities of 1 fmol/cm2 and 10 fmol/cm2 were required to modulate cell spreading and focal adhesion on surface-modified glass substrates.124 However, the effects of peptide concentration on cellular response are highly cellspecific. In another study using similar cells and ECM-based peptides incorporated into non-adhesive hydrogels, the required densities to trigger adhesion, spreading and formation of focal contact was 1000 fmol/cm2 122 indicating that not only the peptide density but also the hydrophilic/hydrophobic nature of the material (and perhaps its mechanical properties) plays an important role. This can be highly significant when working with polysaccharide-based hydrogels, which are highly hydrophilic and can hydrogen-bond with peptide side chains.

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275

Tissue engineering approaches targeting cell signaling

The introduction of ECM proteins and ECM-derived peptides is inescapable when designing tissue-engineered strategies aimed at mimicking aspects of the ECM. Since cell adhesion is often the first step in a cascade of events influencing cell fate leading to differentiation, migration, proliferation or apoptosis, and the adhesive properties of the ECM have been demonstrated in vitro, various tissue-engineering strategies have tried to mimic the celladhesive nature of ECM by surface and bulk modification of biomaterials to enhance cell adhesion and spreading.1,88,125,126 Due to its simplicity and its abundance within the glycoproteins such as fibronectin, collagens, laminins, tenascin, thrombospondin and vitronectin6,47, the RGD motif is usually the first one to be considered. Peptides containing the RGD motif have been introduced on scaffolds derived from alginate,127–131 chitosan,54,132–135 agarose,58,59 hyaluronan136–138 and dextran.69,70,87 Other peptides comprising amino sequences such as IKVAV and YIGSR have been immobilized on dextran to enhance neuron adhesion69 and amino acid sequences such as KQAGDV, ATLQLQEGRLHFXFDLGKGR, GEFYFDLRLKGDKY to promote fibroblast adhesion on chitosan.93 One of the main advantages of using peptides to promote cell adhesion is the specificity of the adhesion mechanism being involved. Many peptides such as those containing the RGD motif promote the cell attachment in a specific manner because the motif is recognized by adhesion cell-membrane receptors such as integrins. Competitive inhibition of cell attachment using soluble peptides showed that the interaction of the cells with the peptidecontaining materials was ligand specific.54,129 While most studies use peptides comprising the RGD motifs aiming for ligand-specific but non-cell-selective cell adhesion, identifying integrin-binding motifs also makes it possible to create cell-selective substrates. Previously, IKVAV-modified substrates were reported to promote substantial neuron cell adhesion and minimal fibroblast and glial cell adhesion139 whereas KHIFSDDSSE-modified surfaces selectively enhanced cell adhesion of astrocytes but not of fibroblasts.107 As ECM components are involved not only in cell adhesion but also in other important morphogenesis events, researchers have been trying to mimic cell-glycoprotein interactions. In neuronal tissue engineering, many studies have functionalized surfaces to promote cell adhesion and neurite growth on biomaterials using ECM-derived peptides from tenascin-c,101,140 fibronectin141 and laminins.106,108,142,143 The most commonly studied ECM-derived peptides include RGD,58,92,107,130,133,137,139,141,142,144,145 laminin-derived YIGSR92,95, 96,107,121,142,145–147 and laminin-derived IKVAV.92,95,96,107,121,139, 142,147 In some cases, RGD and YIGSR were found to promote cellular adhesion and IKVAV to promote neurite outgrowth.139,142,148 Previously our group demonstrated

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that GYIGSR-modified surfaces led to greater cell adhesion and SIKVAVmodified surfaces promoted greater neurite outgrowth;148 surfaces modified with both GYIGSR and SIKVAV had a synergistic effect on the neurite outgrowth.121 Moreover, surfaces modified with extended laminin-derived oligopeptides, CDPGYIGSR and CQAASIKVAV, which better maintain the native conformation of the active sites, promoted greater cellular response than those modified with the respective minimal peptide sequences.95 Despite the presence of other integrin-binding motifs along the fibronectin backbone, such as LDV, REDV and PRRARV and the synergistic integrin binding of RGD (FIII-10) and PHSRN (FIII-9),149–152 most neurological studies have focused on the RGD motif found within the tenth FN-III repeat. Interestingly RGD can differentially mediate neuronal growth depending on the cell type: it promotes cell adhesion for most cells and neurite outgrowth for certain neuronal cells.104,108,153–156 This is probably related to the diversity of RGD-binding integrins. The RGD sequences have been found to bind to many integrin receptors such as: α3β1, α5β1, α8β1, αvβ1, αvβ3, αvβ6, α6β1. Cell migration and neurite outgrowth are known to be guided by chemotactic and haptotactic cues. This phenomenon has been observed previously in vitro with growth factor gradients157 and was also reported with immobilized RGD and IKVAV gradients. Cells migrated along linear gradients of RGDS peptides immobilized within PEG-based hydrogels,158 whereas linear gradients of IKVAV provided guidance to the neurite outgrowth.159 Affinity between synthetic peptides and integrins plays a key role on the intensity of the cellular response towards the modified substrate. To enhance mimicking of binding domain configuration, cyclic peptides have been investigated and have shown greater affinity to the cell receptors than their linear equivalents, thus increasing cell adhesion.143 Functionalization of peptides was enhanced further by Renner, Moroder and coworkers who designed a photo-switchable cyclic peptide by synthesizing heptapeptide containing RGD with 4(aminomethyl) phenylazobenzoic acid.160,161 Irradiation with a laser flipflops the azobenzene conformation, altering binding affinity of the peptide from a low to a high level. The use of this type of peptide has potential utility to control spatial organization of peptides and ultimately of cells in 3-D scaffolds such as hydrogels. Using an approach similar to Luo and Shoichet, who guided neurite outgrowth through biochemical peptide channels within agarose hydrogels,58 light-switchable peptides could be used to design peptide channels within hydrogels that could change geometry and orientation spatially and temporally. The use of stem cells is gaining greater attention in tissue engineering strategies. Since peptides can mimic aspects of the ECM to promote cellular behavior such as cell adhesion and proliferation, interest has arisen for the immobilization of peptides on scaffolds to promote stem cell differentiation.

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Alsberg et al. reported the use of RGD-based peptides immobilized on alginate to control osteoblast differentiation and regenerate bone in vivo.127 Similar differentiation behavior was observed when myoblasts when cultured on similar substrate.129 The use of peptides to promote targeted cell signaling can be advanced beyond mimicking cell-proteoglycan interactions. More recently, scaffolds are being designed with both ECM peptides and growth factors. It is known that most growth factors are sequestered by the ECM and then presented to cells. Thus the latest biomimetic scaffolds are being designed not only with immobilized peptides but with immobilized proteins/growth factors. This has opened up a new way of examining cell behavior and questioning the dogma of growth factor internalization for cellular response. Thus the mechanism of signal transduction can be re-examined with these scaffolds which can also more fully stimulate cellular response. The same chemical modification techniques designed for peptides can be adapted for growth factors. Alternatively, new methods have also been pursued to ensure bioactivity of the growth factor is maintained. Growth factors have been immobilized on a PEG layer, allowing appropriate presentation of the growth factor to the cell.72 Growth factors have also been immobilized as concentration gradients and used to guide cell growth.157,162 Overall, there have been significant advances in peptide-modification of biomaterials and particularly of polysaccharides which are inherently nonadhesive to cells. The almost ‘blank’ background of the polysaccharide allows cell interactions to be dialed into the scaffold design through careful control of the peptide immobilized and the physical and mechanical properties of the scaffold. There are several polysaccharides and numerous peptides that have been investigated, with each chosen based on a set of design criteria dictated by the tissue into which the scaffold will be implanted and the cells of greatest interest. While peptide modification itself cannot create a biomimetic tissue analog, the peptide plays an important role in defining the cellular microenvironment, which is key to cell fate.

9.7

References

1 Lutolf M P and Hubbell J A, Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering, Nat Biotechnol, 2005, 23, 47–55. 2 Huhtala M T, Pentikainen O T and Johnson M S, A dimeric ternary complex of FGFR [correction of FGFR1], heparin and FGF-1 leads to an ‘electrostatic sandwich’ model for heparin binding, Structure, 1999, 7(6), 699–709. 3 Iozzo R V, Matrix proteoglycans: from molecular design to cellular function, Annu Rev Biochem, 1998, 67, 609–52. 4 Frevert C W, Kinsella M G, Vathanaprida C, Goodman R B, Baskin D G, Proudfoot A, Wells T N, Wight T N and Martin T R, Binding of interleukin-8 to heparan sulfate

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5 6 7

8 9

10 11 12 13 14 15 16 17

18

19

20

21

22

Natural-based polymers for biomedical applications and chondroitin sulfate in lung tissue, Am J Respir Cell Mol Biol, 2003, 28(4), 464–72. Stringer S E, The role of heparan sulphate proteoglycans in angiogenesis, Biochem Soc Trans, 2006, 34(Pt 3), 451–3. Rosso F, Giordano A, Barbarisi M and Barbarisi A, From cell-ECM interactions to tissue engineering, J Cell Physiol, 2004, 199, 174–80. Moscatelli D, Metabolism of receptor-bound and matrix-bound basic fibroblast growth factor by bovine capillary endothelial cells, J Cell Biol, 1988, 107(2), 753– 9. Bosman F T and Stamenkovic I, Functional structure and composition of the extracellular matrix, J Pathol, 2003, 200, 423–28. Lodish H, Scott M, Matsudaira P, Darnell J, Zipursky L, Kaiser C, Berk A and Krieger M, Molecular Cell Biology, 5th edition ed., W H Freeman: New York, 2003. Giancotti F G and Ruoslahti E, Integrin signaling, Science, 1999, 285, 1028– 1032. Miranti C K and Brugge J S, Sensing the environment: a historical perspective on integrin signal transduction, Nat Cell Biol, 2002, 4, E83–90. Albelda S M and Buck C A, Integrins and other cell-adhesion molecules, Faseb Journal, 1990, 4(11), 2868–80. Katsumi A, Orr A W, Tzima E and Schwartz M A, Integrins in mechanotransduction, J Biol Chem, 2004, 279(13), 12001–4. Ingber D, Integrins as mechanochemical transducers, Curr Opin Cell Biol, 1991, 3(5), 841–8. Juliano R L and Haskill S, Signal transduction from the extracellular-matrix, Journal of Cell Biology, 1993, 120(3), 577–85. Millward-Sadler S J, Salter D M, Integrin-dependent signal cascades in chondrocyte mechanotransduction, Ann Biomed Eng, 2004, 32(3), 435–46. Millward-Sadler S J, Wright M O, Davies L W, Nuki G and Salter D M, Mechanotransduction via integrins and interleukin-4 results in altered aggrecan and matrix metalloproteinase 3 gene expression in normal, but not osteoarthritic, human articular chondrocytes, Arthritis Rheum, 2000, 43(9), 2091–9. Gigant-Huselstein C, Hubert P, Dumas D, Dellacherie E, Netter P, Payan E and Stoltz J F, Expression of adhesion molecules and collagen on rat chondrocyte seeded into alginate and hyaluronate based 3D biosystems. Influence of mechanical stresses, Biorheology, 2004, 41(3–4), 423–31. Liedert A, Kaspar D, Blakytny R, Claes L and Ignatius A, Signal transduction pathways involved in mechanotransduction in bone cells, Biochem Biophys Res Commun, 2006, 349(1), 1–5. Ignatius A, Blessing H, Liedert A, Schmidt C, Neidlinger-Wilke C, Kaspar D, Friemert B and Claes L, Tissue engineering of bone: effects of mechanical strain on osteoblastic cells in type I collagen matrices, Biomaterials, 2005, 26(3), 311– 18. Kaspar D, Seidl W, Neidlinger-Wilke C, Beck A, Claes L and Ignatius A, Proliferation of human-derived osteoblast-like cells depends on the cycle number and frequency of uniaxial strain, J Biomech, 2002, 35(7), 873–80. Yang G, Crawford R C and Wang J H, Proliferation and collagen production of human patellar tendon fibroblasts in response to cyclic uniaxial stretching in serumfree conditions, J Biomech, 2004, 37(10), 1543–50.

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Part III Biodegradable scaffolds for tissue regeneration

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10 Scaffolds based on hyaluronan derivatives in biomedical applications E. T O G N A N A, Fidia Advanced Biopolymers s.r.l., Italy

10.1

Introduction

Hyaluronan is a glycosaminoglycan ubiquitously distributed in the extracellular space especially in the extracellular matrix. It is a linear polymer consisting of a regular repeating sequence of non-sulphated disaccharide units, glucuronic acid and N-acetyl-glucosamine. Its molecular weight varies from 4000 to 8 × 106 Da with the molecular structure basically unmodified by evolution, which underlies its importance in the physiological environment. It has high capacity for holding water and elevated viscoelastic properties. Hyaluronan, far from being a simple space filler, plays different biological roles depending on its physical and chemical properties. Hyaluronic acid has a direct effect on cell behaviour due to specific interactions with cell surface receptors. In the last few years the use of hyaluronan, in a highly purified form, has become a common practice in medicine. This fundamental polysaccharide is successfully used as a viscosupplementing agent into articular joints and as an ophthalmic substance. Its most advantageous effects appear to be related to improved wound healing in injured tissue. For many reasons, it is reasonable to consider hyaluronan as a promising biomaterial in biomedical applications. In the following section of this chapter is an overview of Hyaluronan properties and the possibilities its modified chemistry could give. The chapter then focuses on tissue engineering strategies that hyaluronan scaffolds allow us to pursue and their relevance in biomedical application. Finally, the last part of the chapter reports future trends in regenerative medicine and suggests a few other sources of relevant information.

10.2

Hyaluronan

Hyaluronan is a major carbohydrate of the extracellular matrix and can be found in skin, joints, eyes and most of the organs and tissues. Hyaluronan is present in all soft tissues of higher organisms and, in particularly high concentrations, in the extracellular matrix of articular cartilage and in the 291 © 2008, Woodhead Publishing Limited

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mesenchyme of the developing embryo (Toole, 2001). It is a linear polymer consisting of a regular repeating sequence of non-sulphated disaccharide units, glucuronic acid and N-acetyl-glucosamine with a molecular weight varying from 4000 to 8 × 106 Da. Remarkably, hyaluronan is basically unmodified by evolution that underlies its importance in the physiological environment. Far from being a simple space filler, hyaluronan plays different biological roles depending on its physical and chemical properties (Liao et al., 2005; Brown, 2004). It interacts with binding proteins, proteoglycans and active molecules, such as growth factors which influence matrix structure, water balance, lubrication and cell interaction. Hyaluronan can influence cell movement by its ability to alter the osmotic pressure leading to the formation of hydrated pathways. This wide range of functional activities may seem surprising given hyaluronan’s simple structure, but in fact this results from the large number of hyaluronan-binding proteins (often termed hyaladherins) that exhibit significant differences in their tissue expression, cellular localization, specificity, affinity and regulation (Toole, 2001). It is clear that the spatial and temporal expression of hyaladherins can modulate the hyaluronan concentration in the environment therefore regulating tissue remodelling. Besides all of these properties, there are indications that hyaluronan is a scavenger molecule for hydroxyl radicals with protein exclusion properties thus offering protection to cells and extracellular matrix molecules against free radical and proteolytic damage (Abatangelo and O’Regan, 1995; Presti and Scott, 1994). The importance of hyaluronan depicted so far in the physiology of the cellular environment is based both on its chemical and physical properties. To further stress the importance of this multi-functional molecule, the role of hyaluronan breakdown products has to be mentioned. It has been reported that oligomers of hyaluronic acid can have pharmacological activity. It has been shown that hyaluronic acid bonded to a substrate exhibits a size-dependent stimulation of chondrogenic differentiation (Kujawa et al., 1986). Other authors have documented the effect of hyaluronan oligomers on vascularization and their angiogenic or anti-angiogenic properties depending on the oligomer size (Deed et al., 1997). Hyaluronan is a major component of the extracellular matrix of embryonic mesenchymal tissues; thus the use of hyaluronan-based scaffolds will help create a milieu supportive for tissue remodelling (Nathanson, 1990; Toole, 2001). Most cells in the body have the capability to synthesize hyaluronan during implicating its function in several fundamental biological processes. It is indeed generally accepted that hyaluronan is associated with the tissue repair process (McDonald and Camenisch, 2002). In recent years the use of hyaluronan, in its highly purified form, has become a common practice in medicine. The unique biophysical properties of hyaluronic acid are manifested in its mechanical function in the synovial fluid, where it confers to the connective tissues the ability to resist compressive

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forces. Not surprisingly, intra-articular administration of hyaluronan is widely used to relieve pain and improve joint mobility in the non-surgical treatment of osteoarthritis (Abatangelo and O’Reagan, 1995; Fernández López and Ruano-Ravina, 2006). The application of hyaluronan in ophthalmology can be cited as another typical medical application (Goa and Belfield, 1994; Menzel and Farr, 1998). It has been hypothesized that hyaluronan-based matrices mimic the embryonic hyaluronic-acid-rich environment supportive for progenitor cell development (Caplan, 2000, 2003). The possibility to elaborate this natural polymer in different physical forms gives the opportunity to translate tissue engineering strategies in clinical practice, providing a biomaterial that induces and modulates the sequence of events that lead to regeneration of damaged tissues.

10.3

Hyaluronan-based scaffolds for biomedical applications

Tissue-engineering strategies utilize specific combinations of cells, scaffolds and bioactive factors with the aim to functionally substitute damaged tissue (Langer and Vacanti, 1993). This field will potentially neutralize drawbacks and limitations of organ transplantations and the use of mechanical devices, such as prostheses. The forces in these recent years that have pushed forward tissue engineering applications were improved techniques in cell culture and increased knowledge in biomaterials sciences. The primary goal of the tissue engineering approach to tissue defects is to provide an engineered construct with the emphasis on regeneration rather than repair. The ideal scaffold should provide an immediate support to cells and have mechanical properties compatible with those of the tissue being repaired. Gradually then the material has to degrade, as the cells begin secreting their own extracellular matrix, thus allowing for an optimal integration between newly-formed and existing tissue (Bell, 1995). Moreover, the material breakdown products should not be toxic and even better possess a biological activity so as to provide informational cues, in this way mimicking the native extracellular environment. In cartilage tissue engineering for instance, the scaffold must promote chondrocyte attachment and in vitro proliferation, as well as favour the expression and maintenance of a cell-differentiated phenotype (Benya and Shaffer, 1982). For the development of a suitable polymer scaffold, it is essential to consider that chondrocytes must organize three-dimensionally to stimulate the synthesis of extracellular matrix molecules and prevent the loss of cell phenotype for a proper cartilage function once implanted (Brun et al., 1999; Girotto et al., 2003; Sharma and Elisseeff, 2004). As so often happens, the best solution is to follow nature’s direction during the tissue formation process (Bullard et al., 2003; Yannas et al.,

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2007). This indeed is a challenging task that involves a multidisciplinary approach. There is the necessity to understand the biology of tissue formation during embryogenesis, to possess material science knowledge to mimic the consistency and architecture of tissues in the adult organism, and to understand tissue mechanical properties and their alterations due to the physiological activity. Moreover, the complexity of cellular biology has to be added to this equation when interactions between cells and extracellular matrix is considered (Ulrich-Vinther et al., 2003). In fact, the ECM is not only acting as a structural support but is involved in a reciprocal informational exchange with the cellular component (Cattaruzza and Perris, 2005; Berrier and Yamada, 2007). As early as the first cellular division occuring during the embryonic development, cells begin to synthesize and secrete many kinds of molecules that form a particular environment providing both a scaffold and a guidance during development and throughout adult life. Although the recapitulation of such embryonic events represents a major goal for tissue engineering, it has to be kept in mind that the adult environment is quite different from the one present during the embryonic stage, meaning precisely that tissues respond differently to specific informational cues (Caplan, 2000). All this is performed through complicated multi-step differentiation cascades implying fine cross-talk between cells and ECM with the synthesis of sitespecific specialized molecules (Toole, 2001; Hubbell, 2003). Therefore, the appropriate scaffold should be designed to play this multi-functional role and not be used as a mere vehicle for reparative cells or growth factors. The scaffold has to change its features as the regenerative process takes place, starting with tissue induction in the early events and then supporting tissue formation during differentiation. With this notion in mind the materials used in tissue engineering have to be biocompatible and biodegradable. Physical properties have to match those of the tissue being replaced, promoting cell attachment and ultimately possessing the capability of being remodelled by tissue specific cells (Frenkel and Di Cesare, 2004). Our approach to these issues is to rely on hyaluronan-based biomaterials. This natural molecule appears to be an ideal candidate for tissue engineering strategies because of its dual role, both structural and informational. The greatest barrier for the successful use of this polymer in tissue engineering is the fact that hyaluronan, in its purified form, has certain characteristics that limit its use as a biomaterial. Water solubility, rapid resorption and short residence time in the tissue along with its poor ductility hamper its possible applications. Hyaluronic acid exists only as an aqueous gel which is rapidly degraded upon application. Cross-linking and coupling reactions are two ways to obtain a modified, stable form of hyaluronan (Soranzo et al., 2004) (Fig. 10.1). The first involves the modification of specific functional groups of the molecule by subjecting them to chemical reaction such as esterification or amidation, while the second involves the

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–CH2OH ✓ Carboxyilated ✓ Crosslinks

OH

CH2OH

HO

HO O

O NHCOCH3

–NHCOCH3 ✓ De-N-Acetilation ➣ N-sulphation ➣ Amides

O

ONa

–COO– ✓ Esters ✓ Amides ✓ Auto-cross-linked ✓ Cross-linked with bridge molecules

10.1 Reactive groups that can undergo chemical modifications on the native hyaluronic acid backbone.

creation of reticules between the molecule polymeric chain through condensation of specific functional groups or alternatively through the formation of atomic bridges. It is important to note that the esterification reaction may involve all or some of the carboxyl groups of hyaluronic acid giving rise to compounds with quite different physicochemical properties. An esterification involving more than 50% of the carboxyl groups tend to drastically reduce water solubility, leading to a virtually water-insoluble polymer for esterification degrees above 70%. With the application of these techniques, it is possible to improve hyaluronan processability. Hyaluronan derivatives can be extruded to produce fibres or membranes, lyophilized to obtain sponges or processed to obtain microspheres (Fig. 10.2). Moreover, fibres can be carded into gauzes, ropes or non-woven structured materials. Combining hyaluronan properties with the possibility of having it in a more stable and workable form makes hyaluronan derivatives one of the most promising materials in the biomedical sciences today. HYAFF® is the class of hyaluronan derivatives obtained by a coupling reaction. The approach behind the creation of these biopolymers was to improve hyaluronan stability by esterifying the glucuronic acid group with different types of alcohols. This chemical modification acts by reducing the hydrophilic component and increasing the hydrophobicity of the parental molecule at the same time. By changing the type and the percentage of esterification, it is possible to create a broad variety of polymers with different consistencies and residence times (Campoccia et al., 1998).

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Hyaluronic acid

HYAFF®

Extrusions

Membrane

Thread

Entrapment

Guide

Twisting Combi-twisting

Gauze

Carding Needle-carding

Textiles Non-woven

Lyophilization

Drying Extraction Evaporation

Granulation

Sponge

Microspheres

Granules

Perforation

Perforated membrane

10.2 Chart exemplifying the wide array of possibilities that HYAFF-based biomaterials can give in regenerative medicine and biomedical applications. Structural characteristics of the scaffold are fundamental in tissue reconstruction. © 2008, Woodhead Publishing Limited

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Chemical modifications

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HYAFF ® 11 is the most characterized biopolymer from both the physiochemical and the biological point of view (Milella et al., 2002; Soranzo et al., 2004). It is produced starting from hyaluronan of about 200 KDa and represents the total benzyl ester of the parent molecule. This biopolymer can be processed into stable configurations to produce a variety of biodegradable structures with different physical forms and in vivo residence times. Extensive biocompatibility studies have demonstrated the safety of HYAFF®11 and its ability to be reabsorbed in the absence of an inflammatory response (Campoccia et al., 1998). ACP® is instead an ester obtained through a condensation reaction resulting in the formation of inter- and intramolecular ester bonds (Fig. 10.3). During the reaction a predetermined percentage of carboxyl groups is esterified with hydroxyl groups of the same molecule, thus forming a mixture of lactones and intermolecular ester bonds. Thermodynamic calculation shows that the crosslinks preferably involve groups on different chains, thus giving rise to intermolecular bonds. It is relevant to note that no foreign bridge molecules are attached to the Hyaluronan chain, thus ensuring only the liberation of parental hyaluronan upon degradation. ACP®, compared with the native hyaluronic acid, presents significantly improved viscoelastic properties (Mensitieri et al., 1996). The absence of foreign bridge molecules ensures the release of native hyaluronic acid only during degradation, while the autocrosslinking process improves the viscoelastic properties of the gel compared with unmodified hyaluronan solutions of the same molecular weight. Because of its very high hydrophilicity, in vitro degradation is quite rapid. It has been hypothesized that hyaluronan-based matrices mimic the embryonic hyaluronic-acid-rich environment supportive of progenitor development. When used as a cell carrier in adult organisms, HYAFF® and ACP®-based matrices

O

O

O

NHCOCH3

OH

O

O OH

O O

OH O

NHCOCH3

OH

O

O

O

O OH

O O

OH O

10.3 Sketch illustrating (in the black outlined box) the groups that in ACP are involved in the autocross links.

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tend to promote the recapitulation of those events that facilitate tissue repair (Solchaga et al., 1999; Caplan, 2000, Caplan, 2003).

10.4

Clinical applications

10.4.1 Post-operative adhesions Post-operative adhesions occur after most peritoneal surgical procedures and can result in serious complications, including intestinal obstruction, pain and infertility in women (Holmdahl and Risberg, 1997). Adhesions are abnormal attachments between tissues and organs and may be congenital or acquired. The development of acquired adhesions is a generalized phenomenon in response to trauma to the peritoneum that could arise from inflammation or following surgery. Most surgical gynecological/obstetrical procedures are commonly associated with adhesion formation, such as ovarian cystectomy, myomectomy, total abdominal hysterectomy, endometriosis and adhesiolysis. Adhesions result from the normal peritoneal wound healing response and develop in the first five to seven days after injury. Adhesion formation begins with coagulation which initiates a cascade of events resulting in the build up of fibrin gel matrix. If not removed, the fibrin gel matrix serves as the progenitor to adhesions by forming a band or bridge when two peritoneal surfaces, coated with it, are apposed. In practice the band or bridge becomes the basis for the organization of an adhesion. Protective fibrinolytic enzyme systems of the peritoneum, such as the plasmin system, could remove the fibrin gel matrix. However, surgery dramatically diminishes fibrinolytic activity (Di Zerega, 1997; Di Zerega and Campeau, 2001). The major strategies for adhesion prevention or reduction are adjusting surgical practice (minimizing the invasiveness and surgical trauma) and applying adjuvants. One of the most promising methods of preventing adhesions due to surgery involves placement of a mechanical barrier between raw tissue surfaces that keeps those surfaces apart. The ideal barrier method for adhesion prevention should be safe, proven effective, easy to use in both laparoscopic procedures and laparotomy, absorbable, noninflammatory, should not require sutures, should not potentiate infection and should not interfere with wound healing (Harris et al., 1995). One particularly promising biopolymer that effectively acts as a barrier is hyaluronic acid (Nappi et al., 2007). Hyaluronic acid has been experimentally shown to reduce postoperative adhesion formation after abdomino-pelvic and orthopaedic surgery, the antiadhesive effects depending on the molecular weight as well as the concentration of the preparation. The results were variable because unmodified hyaluronan was subject to rapid degradation and was cleared from the site of administration within hours. On the other hand, a modified form of hyaluronic acid has been reported to reduce incidence and severity of adhesions. A new class of hyaluronic acid derivative, the auto-crosslinked polysaccharides series (ACP®) has been © 2008, Woodhead Publishing Limited

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developed with the aim of increasing the viscosity and the residence time of the gel upon application (De Iaco et al., 1998 and 2001). Hyalobarrier® gel is made of a highly viscous biocompatible and biodegradable aqueous gel of an autocross-linked hyaluronic-acid derivative with improved viscoelastic properties compared with unmodified hyaluronan solutions of the same molecular weight. Hyalobarrier® gel has a high adhesivity and a prolonged residence time. As no foreign molecules are used to create the cross-link, the gel upon degradation is de-esterified and liberates hyaluronic acid, a naturally occurring molecule that has a biological role and is eliminated through known metabolic pathways. Preclinical trials in animal models have shown that Hyalobarrier® gel reduces the incidence and severity of postoperative adhesions (De Iaco et al., 1998; Belluco et al., 2001; De Iaco et al., 2001; Pucciarelli et al., 2003). Moreover, preliminary clinical studies in hysteroscopic surgery as well as laparotomic and laparoscopic myomectomy have suggested that Hyalobarrier® gel may reduce the incidence and severity of postoperative adhesions in pelvic surgery (Acunzo et al., 2003; De Iaco et al., 2003; Pellicano et al., 2003; Carta et al., 2004; Guida et al., 2004) with improvement in the pregnancy rate in infertile patients who were submitted to laparoscopic myomectomy (Pellicano et al., 2005). More recently, the favourable safety profile and the efficaciousness of this ACP®-based gel has been confirmed in a blind, controlled and multicentre study (Mais et al., 2006). In conclusion, on the basis of its efficacy, safety and ease of use, Hyalobarrier® gel can be considered today as one of the most valuable site-specific anti-adhesion barriers for abdomino-pelvic surgery. Exploiting similar technology is also Hyaloglide®. This is a highly viscous ACP-based absorbable gel providing a hyaluronan surgical barrier for the prevention of fibrosis after tendon and nerve surgery (Brunelli et al., 2005). Following in vivo trials, Hyaloglide® was tested in a controlled clinical study to investigate its safety and efficacy after microsurgical reconstruction of peripheral nerves of the hand (Atzei et al., 2007). No safety concerns were raised and no adverse effects on either wound closure or nerve repair were observed at follow-up visits in the treatment group, and the applicability of this hyaluronan-based gel was judged as favourable in all cases. Therefore, the application of Hyaloglide® was shown to improve recovery of sensitivity and pain following microsurgical repair of peripheral nerves by limiting formation of perineural adhesions and favoring regeneration of nervous tissue, creating a more favourable environment for the healing of the lesion.

10.4.2 Adhesion formation following sinus and otological surgeries The widespread practice of endoscopic surgery in a multitude of disciplines has induced commercial industry to develop novel bioabsorbable materials,

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frequently used in the paranasal sinuses to control bleeding, promote tissue regeneration, improve healing, or prevent scarring. Such materials have been employed in place of traditional tampons or petroleum gauze packing with the ultimate intent of improving patient comfort and reducing the incidence of complications associated with nonabsorbable packing (e.g. mucosal injury, pack dislodgement and aspiration, obstructive sleep apnea secondary to nasal obstruction, toxic shock syndrome, foreign body granuloma, myospherulosis, and patient discomfort). Several different types of topical packing agents have been investigated including gelatin, collagen, paraffin gauze, hydroxylated polyvinyl-acetal polymers and several hyaluronan polymers (Chandra and Kern, 2004; Incesulu and Häusler, 2007). Belonging to this latter family, Merogel is a product based on hyaluronic acid that is used as an adjunct in endoscopic sinus surgery and middle ear surgery as an absorbable material intended to assist in wound healing through the minimization of scarring and adhesions and the encouragement of epithelialization. The product is a woven form of the benzyl ester of hyaluronic acid, which is intended to absorb significant quantities of fluid in the wound site working over a period of two weeks. The general biocompatibility of this hyaluronic derivative in these kinds of application has been established over long periods of time (Zhong et al., 1994; Campoccia et al., 1996; Campoccia et al., 1998). The clinical use of Merogel proved its safety, well-acceptance and tolerability with the significant advantage of its being resorbable. Moreover, Merogel may favour improved healing in patients undergoing nasal surgery and reducing adhesion formation (Miller et al., 2003; Catalano and Roffman, 2003; Xu et al., 2003; Wormald et al., 2006; Franklin and Wright, 2007). A summary of possible configurations in these types of application is detailed in Table 10.1. Another important application of Hyaluronan derivatives is in the field of wound healing (Chen and Abatangelo, 1999). Skin is the body’s largest Table 10.1 Possible hyaluronan-based hydrogels used in adhesion prevention Product

Description

Indication

Merogel Merogel Injectable Meropack

Hyaluronan biodegradable stents (gel, pad)

Improved healing after sinus surgery

Epifilm

Hyaluronan biodegradable implant (film, pad)

Surgical adjunct in otological surgery

Hyalobarrier

Hyaluronan surgical barrier (gel)

Reduction of post-surgical adhesions in abdomino-pelvic surgery

Hyaloglide

Hyaluronan surgical barrier

Prevention of fibrosis after tendon and nerve surgery

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organ. It is comprised of an epithelium, the epidermis and a connective tissue matrix, the dermis. The skin’s primary function is to act as a protective barrier between the body and the external environment. The skin blocks penetration of micro-organisms, absorbs and blocks radiation, and inhibits the loss of water from the body. The skin also is involved in thermoregulation (prevents heat loss and allows for rapid cooling through evaporation of sweat) and functions in immunologic surveillance by assisting in the presentation of antigens to immune cells. Therefore extensive wounds require a barrier protection to prevent infection and desiccation and cell guidance by dermal elements to maximize healing. Hyalomatrix® PA is a bi-layered, sterile, flexible and conformable wound dressing which acts as a dermal substitute in the treatment of acute wounds. Hyalomatrix® PA is indicated as a dermal substitute for the immediate coverage, following surgical excision and prior to grafting, of deep burns. It is also indicated in the treatment of partial- or full-thickness post-traumatic, post-surgical wounds. The wound contact layer is constituted by an absorbent, biodegradable non-woven pad entirely composed of HYAFF®. The HYAFF® layer is physically coupled with a transparent and flexible film of a medical grade synthetic elastomer that acts as a semi-permeable barrier towards external contaminants. As Hyalomatrix® PA is applied on the wound bed, the biomaterial wound contact layer provides a three-dimensional scaffold able to be colonized by fibroblasts and onto which extracellular matrix components are regularly laid down, favouring an ordered reconstruction of the dermal tissue. The transparent elastomeric film, characterized by a vapour transmission rate comparable to that of normal skin, avoids excessive body fluid loss and maintain a moist wound environment. The transparency of the film allows also for a continuous wound inspection without need of its removal. This dermal substitute conforms well to wound edges also because it has the advantage of being cut to suit the shape of the wound. On the basis of its clinical use, Hyalomatrix® PA represents a new step towards a potentially faster wound closure which could potentially improve aesthetic outcomes (Price et al., 2006; Gravante et al., 2007) (Table 10.2).

10.4.3 Hyalonect® The periosteum or periosteal membrane is a continuous composite fibroelastic covering membrane of the bone to which it is intimately linked. Although the bone cortex is the main beneficiary of the principal anatomical and physiological functions of the periosteal membrane, the behaviour of the entire bone remains closely influenced by periosteal activity. These principal functions are related to the cortical blood supply, osteogenesis, and muscle and ligament attachments. Through its elastic and contractile nature, it participates in the maintenance of bone shape, and plays an important role in

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Table 10.2 Possible hyaluronan-based scaffolds to be used in chronic wound management Product

Description

Indication

Laserskin and Hyalograft 3D Autograft

Bioengineered hyaluronan dermo-epidermal grafts

Hard to heal chronic wounds

Hyalomatrix

Dermal substitute

Acute wounds (consequent to burns and trauma)

Hyalofill Hyalogran Jaloskin

Biological wound dressings

Management of chronic and acute wounds

Note: Laserskin Autograft and Hyalograft 3D Autograft are approaches based on tissue engineering and therefore use cell-seeded scaffold; the other options rely on physico-chemical properties of the biomaterial

metabolic ionic exchange and physiologic distribution of electro-chemical potential differences across its membranous structure. It has also been suggested that the periosteum may have its own specific proprioceptive property. The role of periosteum in the blood supply of bones has been studied by several investigators. Its osteogenic potential, particularly its important role in fracture repair, is also well recognized. Also the importance of periosteum in the biomechanics of bone fracture has been studied in detail, demonstrating that the presence of periosteum contributes to increased resistance to the fracture of bone. It is well known that preservation of both structural and functional integrity of periosteum may play an important role in rapid bone regeneration, since it guarantees cortical blood supply and an osteogenic cell source to the underlying bony tissue. It is therefore fundamental to properly reconstruct the periosteal membrane when a bone injury results in either a partial or total damage to its structure. A periosteum replacement patch may legitimately be considered fundamental to encouraging tissue regeneration of underlying bone graft materials used to treat bone defects as well. By working with different degrees of esterification and solubility, it is possible to manufacture medical devices made of pure hyaluronic acid in a solid form, for example Hyalonect®, which is a surgical mesh composed of HYAFF® fibres. Hyalonect® could be a useful resorbable medical device to avoid possible bone graft material extravasation out of bone defects and guarantee its stability and contact with host tissue. Hyalonect® is a biodegradable and suturable surgical mesh intended for use in reconstructive surgical procedures as an implant material to aid in the natural healing processes of surgical repairs and in tissue containment (Table 10.3). It may be employed to facilitate adhesion at the suture line and/or to stabilize layers of adjacent tissues, thereby allowing physiological tissue repair. The degradation products, namely oligosaccharides,

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Table 10.3 Possible hyaluronan-based scaffolds to be used in musculoskeletal applications Product

Description

Indication

Hyalograft C Autograft

Bio-engineered autologous cartilage graft

Acute cartilage defects

Hyalonect

Biodegradable periosteal patch

Musculoskeletal surgery, periosteal coverage, promotes stability of bone substitutes

Hyaloss Matrix

Adjunct in dental surgery where bone grafting is required

Periodontal surgery, perimplant bone defects, post-extraction bone defects, crest and sinus augmentation

Note: Hyalograft C Autograft is a tissue-engineered cartilage implant while Hyalonect and Hyaloss Matrix are cell-free implants

are implicated in several biological processes including angiogenesis, differentiation and morphogenesis, cell migration and aggregation and development of mesenchymal cells. These biological properties reveal the potential of HYAFF® biopolymers for the development of medical-surgical devices for application in musculoskeletal and orthopaedic surgeries, both as a guide for the newly-forming tissues and as carriers to delivery cells and growth factors. Hyalonect® is a mesh to be used as bone graft containment material and artificial periosteum. Bone grafting is a widely used and universally accepted treatment in orthopaedic practice to deal with problems associated with bone loss, reconstruction, and delayed union or non-union. These bone defects can be large and require the use of bulk graft and/or significant amounts of particulate graft material. Autologous cancellous bone graft remains the most effective grafting material because it provides the three elements required for bone regeneration: osteoconduction, osteoinduction, and osteogenic cells. However, because autogenous grafting is associated with obstacles and complications, including limited quantities of bone for harvest and donorsite morbidity, alternatives have been used in a wide range of orthopaedic pathologic conditions. Grafting substitutes currently available include cancellous and cortical allograft cone, ceramics, demineralized bone matrix paste and putty, bone marrow and composite grafts. In any case, an adequate and functional bone tissue regeneration depends on the interaction of the bone graft material and the host’s mechanical and biological environment, and host-bone graft contact and stability. Experimental and clinical studies have shown that when host-graft interfaces were fixed securely and remain stable, rapid healing and union occurred between the bone graft and host bone. If the host soft-tissue bed is compromised in any way, the

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revascularization and migration of osteoprogenitor cells will be compromised and as a result bone grafts will fail to be incorporated into the host tissue. Hyalonect® could be a useful resorbable medical device to avoid possible bone graft material extravasation out of bone defects and guarantee its stability and contact with host tissue. As a result of its physical and biological properties, it could play an important role in bone restoration and in the repair of damaged or inadequate integumental tissue, such as periosteum, by minimizing fibrous tissue migration into the underlying graft site, favouring revascularization of the graft site to promote nutritional diffusion, cell proliferation and osteogenesis within the bone defect, creating a protected environment for tissue and bone healing.

10.4.4 Dental surgery Another application of hyaluronan-based biomaterials involving both soft and hard-tissue is in dental surgery. Hyaluronic acid’s regenerative powers were discovered as early as 1700. The ‘dentists’ from that time who were experimenting with the first dental transplants saw that freshly-drawn teeth temporarily grafted onto cocks’ combs took and remained perfectly viable. Cocks’ combs are in fact a source of hyaluronic acid even though nowadays for safety and ethical reasons the polymer is obtained via fermentation from bacteria. With today’s far vaster knowledge, technology has made important headway including biomaterial science, as we have seen. HyalossMatrix® was designed as an adjuvant in the surgical application of bone grafts in the treatment of dental defects and is formed entirely from the HYAFF® biopolymer in the form of fibres (Table 10.3). On contact with the patient’s blood or with saline solution, HyalossMatrix® forms a gel almost instantly, thus facilitating the application of bone fragments. HyalossMatrix® is highly versatile because, at room temperature, it is able to form a gel which the dentist can adjust to the desired consistency by adjusting blood or saline volume. The biodegradable and biocompatible gel thus formed can be mixed with bone fragments to form a paste which is then easily moulded to fit the bone defect. HyalossMatrix® has a double function: on one hand its physicochemical characteristics facilitate application of the bone graft to the site of the damage, while on the other hand it acts as a source of information, creating a hyaluronic acid-rich environment in the damaged area, with all the advantages that this offers (as described in previous sections). Hyaluronic acid’s known biological functions on angiogenesis and the behaviour of mesenchymal stem cells as well as on differentiation and morphogenesis enable HyalossMatrix® to favour the physiological processes of tissue repair, thus stimulating the graft to take naturally and bone regeneration to take place. More complex applications of hyaluronan derivatives involve the use of both biomaterials and living material to have combined cell-scaffold constructs

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used to restore damaged tissues. This is a two-stage procedure that involves a first step in which a tissue biopsy is taken to collect cells. Following this first phase of cell extraction and expansion, the scaffold is loaded with a sufficient amount of cells and then, or after a period of in vitro culture, implanted into the patient. The last two examples provided in this brief overview of clinical application of hyaluronan derivatives concern articular cartilage defects and skin substitutes.

10.4.5 Articular cartilage Articular cartilage in adults has limited ability of self-repair related to the absence of vascularization and the presence of few and very specialized cells with low mitotic activity (Solchaga et al., 2000). Cartilage lesions are a common cause of disability, often associated with pain, reduction of joint mobility and loss of function. In recent years several options have become available to treat cartilage lesions. Treatment algorithms have been proposed to choose the appropriate surgical repair technique, based upon clinical symptoms and patient characteristics and expectations as well as the size and depth of the cartilage lesions (Cain and Clancy, 2001). Traditional surgical techniques, that is palliative or reparative options, have shown variable success rates and carry a number of limitations. Given the intrinsic limitations of these techniques, newer surgical approaches have been developed, focused on obtaining a complete regeneration of the hyaline cartilage as well as a complete integration with the surrounding tissues, to restore the normal knee function and to provide durable outcomes. The cell-based approach, known as Autologous Chondrocyte Implantation (ACI), was first introduced in Sweden in 1987. This technique requires an arthroscopic harvesting of a small cartilage biopsy from a non-weight bearing area of the knee and the subsequent transplantation of in vitro expanded autologous chondrocytes to the defect site, beneath a periosteal flap obtained from the tibia in the same surgical session (Brittberg et al., 1994). Recent evidence from prospective, controlled clinical studies demonstrate that ACI provides a higher and longer-lasting clinical benefit compared with other conventional surgical options for cartilage repair, supporting the hypothesis that cell-based treatment is ‘regenerative’, not only ‘reparative’ (Peterson et al., 2000). The regeneration of new tissue will lead, in fact, to the formation of a normal and site-specific tissue with optimal integration, therefore fully functionable, whereas repair simply gives rise to scar tissue. Despite the excellent long-term clinical results of ACI, there is considerable interest in improving this technology. In fact, the use of ACI is associated with a number of limitations essentially correlated with the complexity and morbidity of the surgical procedure, as well as the frequent occurrence of periosteal hypertrophy (Brittberg et al., 2003; Peterson et al., 2000). To

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overcome these limitations, tissue engineering techniques have emerged as an innovative field of research with the potential to recreate three-dimensional structures such as cartilage, to be used as a replacement for damaged tissue and to regenerate a functional organ in vivo (Caplan, 2000; Langer and Vacanti, 1993). Chondrocytes can be easily isolated from small biopsies and expanded in vitro, but when cultured on two-dimensional substrate they tend to lose their characteristic phenotype. The maintenance of differentiated phenotype is favoured by the use of a three dimensional scaffold (Benya and Shaffer, 1982). HYAFF®11-based scaffolds not only allow chondrocyte attachment but also provide adequate support to permit the expression of the differentiated chondrocyte phenotype (Grigolo et al., 2002). This well correlates with the findings that hyaluronan chains of 200–400 kDa can induce embryonic mesenchymal stem cells to differentiate into chondrocytes (Kujawa et al., 1986). Therefore these scaffolds have a dual intrinsic function: initially, a structural support with high molecular weight hyaluronan and then informational as small hyaluronan oligomers are released during scaffold degradation. The scaffold used in this application consists of a network of 20-µm-thick non-woven fibres with interstices of variable sizes, which constitutes an optimal physical support for the cells (Table 10.3). In vitro studies indicate that chondrocytes adhere, remain viable, and proliferate within the biomaterial. Chondrocytes grow throughout the non-woven scaffold expressing again their original phenotype, which is lost during the expansion phase in monolayer culture (Aigner et al., 1998; Grigolo et al., 2002). Further studies showed that chondrocytes seeded into HYAFF®11-based scaffolds are able to synthesize a hyaline-like matrix, rich in glycosaminoglycans and type II collagen, and positive to metachromatic toluidine blue stain (Campoccia et al., 1998; Brun et al., 1999). In vivo experiments have confirmed these in vitro results. Full-thickness condral defects of the femoral condyle in rabbits treated with autologous chondrocytes cultured on the HYAFF®11 matrix demonstrated a hyaline-like cartilage regeneration (Grigolo et al., 2001; Solchaga et al., 2000). Ultimately, pre-clinical findings supported the use of HYAFF®11 scaffolds for autologous chondrocyte implantations in a clinical setting and led to the development of Hyalograft®C autograft. Hyalograft®C autograft is then defined as a tissue engineered graft composed of autologous chondrocytes grown on a three-dimensional HYAFF®11 scaffold. This innovative approach was first introduced into clinical practice in a number of European countries in 1999 for the treatment of full-thickness cartilage defects and has generated growing interest in the orthopaedic community, with more than 4800 patients treated so far. Reports from randomized controlled trials and clinical cases support the safety and efficacy of Hyalograft®C autograft. Overall, the results indicate that treatment with Hyalograft®C autograft is associated with improvements

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in relief of symptoms, mobility, pain reduction, and quality of repaired cartilage tissue in the majority of patients, with a limited number of adverse events reported. (Marcacci et al., 2005; Nehrer et al., 2006; Tognana et al., 2007).

10.4.6 Skin tissue The importance of skin tissue has already been mentioned. More complex, hard to heal chronic wounds still represent a clinical relevant challenge. A chronic wound is defined as a wound that does not heal in the time expected based upon the patient’s age, co-morbidity, wound location, wound size and wound aetiology. Chronic wounds, including venous ulcer, lower extremity diabetic neuropathic ulcers and pressure ulcers, lead to high rates of morbidity and mortality, diminished quality of life and high healthcare costs. Ideally, conventional wound therapy would result in healing but in the event of nonhealing wounds, more aggressive forms of therapy, such as the application of a skin substitute or the use of growth factors, may be indicated. Recently, tissue engineering strategies have been employed in an attempt to avoid the use of allogeneic (i.e. originating from another individual) skin to re-establish a proper and physiological tissue continuity. Tissue engineered skin substitutes (i.e. human skin equivalents) are products that use living cells (e.g. fibroblast and keratinocyte) in a scaffold of natural, biodegradable or synthetic matrices to foster wound healing. Skin substitutes are generally comprised of epidermal cells, dermal cells or composites (i.e. combinations of dermal and epidermal), and may be used as either temporary or permanent wound coverings. They are indicated for treatment of wounds that have not responded to aggressive, conventional wound therapy. The skin substitutes currently available for clinical use have been developed using different approaches: (a) autologous or allogenic cells cultured on different scaffolds; or (b) acellular biocompatible/biointeractive matrix. Thus different types of skin substitutes are on the market with the intended use to help skin reconstruction, especially when a deep wound occurs. The first epidermal substitute, based on Rheinwald and Green’s keratinocyte culture method and commercially available, was Epicel, manufactured by Genzyme (USA) since 1987. Epicel is produced from epidermal cells harvested from a patient’s skin biopsy, expanded in-vitro to create individual grafts which are then supported by a moisture-retaining petrolatum impregnated gauze dressing. The Epicel culture process utilizes routine mammalian cell culture techniques including an irradiated murine (3T3) cell feeder layer that supports early growth of epidermal keratinocytes. As the keratinocytes proliferate on the culture flask, the feeder layer is displaced from the flask and removed during subsequent culture medium changes. However, it now is broadly accepted that the presence of a scaffold supporting the keratinocyte layers might overcome several problems, such as fragility, handling, contraction,

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and, with an appropriate preparation of the wound bed, improve the take rate, otherwise unpredictable before grafting (Zacchi et al., 1998). To overcome technical problems related to the use of cultured keratinocytes, many workers have developed different biomaterials to function as delivery systems for keratinocytes seeding on the wound site. Fidia Advanced Biopolymers has developed Laserskin ® , a micro-perforated HYAFF-based membrane specifically designed as support for keratinocytes growth and proliferation, for the production of the epidermal substitute Laserskin® autograft. This is an epidermal substitute composed of autologous keratinocytes and indicated in the treatment of chronic wounds (diabetic foot ulcers, vascular ulcers and post-traumatic ulcers) and acute skin wounds (burns and traumatic soft tissue loss) (Campoccia et al., 1998). Moreover, several studies have shown that the presence of a dermal layer is of great importance in the regulation of the growth and differentiation of cultured keratinocytes. Indeed, fibroblasts embedded in their extracellular matrix constitute a permissive and regulatory layer for keratinocytes that otherwise would express a compromise pattern of adhesion molecules and extracellular matrix proteins. Cultured fibroblasts, in common with most cells, grow in two dimensions; therefore three dimensional structures have been used as support. Several matrices (e.g. collagen, glycosaminoglycans and synthetic polymers) could also be exploited to provide such support. Hyalograft®3D is a biodegradable fibroblast delivery system which has been specifically designed to address problems associated with fibroblast grafting. Its non-woven geometry, totally composed of HYAFF®-11, proved to be ideal for fibroblast attachment and proliferation (Campoccia et al., 1998). Therefore, altogether, it would be possible to have epidermal and dermal composites entirely based on hyaluronan technology to be used as permanent wound coverings (Caravaggi et al., 2003; Uccioli 2003; Cavallini, 2007). See Table 10.2 for a list of possible hyaluronan-based scaffolds to be used in wound management.

10.5

Future trends

Tissue engineering has completed impressive achievements but successful regenerative therapies are still not yet available. Hyaluronan presents a variety of multi-functional activity being both a structural and informational molecule. Traditionally, it was thought to be associated with the extracellular matrix, but hyaluronan also has roles inside the cell. Investigation of hyaluronan synthesis and degradation, the identification of new receptors and binding proteins and the elucidation of hyaluronan-dependent signalling pathways keep providing novel insights into the true biological functions of this intriguing polymer. The possibility to elaborate this natural polymer in different physical forms has clearly given the opportunity to develop several devices to be used

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as smart biomaterials in regenerative medicine or for culturing cells and then subsequently implanting tissue engineered based constructs in patients to restore the damaged tissue. The brief overview of evidence presented here hopefully highlighted how hyaluronic acid technology and the tissue engineering approach could improve clinical outcomes of regenerative medicine. Nevertheless much needs to be done to meet clinical challenges ahead. In fact, the ageing population, the high expectations for better quality of life and the changing lifestyle of society leads to the need for improved, more efficient, and fundamentally affordable health care. To this end I believe the further development of nanotechnology and biomaterial sciences could play a pivotal role in improving the regenerative medicine technology platform. What is now called nano-medicine has potential impact on the prevention, early and reliable diagnosis and treatment of diseases. It exploits the improved and often novel physical, chemical, and biological properties of materials at the nanometric scale. Yet, hyaluronan, this long known molecule, with its simple repeating disaccharide structure maintains a central role in present and future biomedical applications.

10.6

Sources of further information and advice

As the first source to have updated and given information in depth about hyaluronic acid and its application, the Glycoforum website (www.glycoforum.gr.jp) is an appropriate tool. To further develop knowledge on tissue engineering and regenerative medicine available is the book edited by Lanza, Langer and Vacanti Principles of Tissue Engineering, now in its third edition. Finally, to have an European overview of the state of the art and progression of tissue engineering along with its strategies and possible applications there are two main sources. Two European projects expressly deal with biomaterial sciences, biomedical applications and tissue engineering but from different perspectives. The first with a strong network is more academic while the second is more industry driven. Both are under the Sixth Framework Programme of the European Union. The first is Expertissues, a network of excellence that brings together the most outstanding scientists in this field throughout the European Union (www.expertissues.org). The other is Systems Approach to Tissue Engineering Products and Processes (STEPS). STEPS has been designed to fill those gaps that presently limit the full exploitation of the multidisciplinary potential of tissue engineering (www.stepsproject.com). A broader view of research activities inside the European Union can be found at the Community Research & Development Information Service portal (CORDIS), http://cordis.europa.eu/en/home.html.

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11 Electrospun elastin and collagen nanofibers and their application as biomaterials R . S A L L A C H and E . C H A I K O F, Emory University/Georgia Institute of Technology, USA

11.1

Introduction

Current pursuits in the discipline of biomedicine, including artificial organs and engineered living tissues, are dependent on the ability to generate novel materials, fabricate or assemble materials into appropriate 2-D or 3-D structures, and precisely tailor material-related properties in order to achieve a desired clinical response (Chaikof et al., 2002). To that end, of profound importance is the development of artificial extracellular matrices (ECM). These structures are integral to the fashioning of microenvironments that are engineered for ideal mechanical and biological performance. It is likely this design will require the mimicry of many, if not all, morphological or physiologic features of native tissues. Decades of research have indeed demonstrated that as our ability to control the physical and biological properties of scaffolding materials improves, the quality of the tissues thus formed is enhanced. More specifically, molecular and supramolecular organization of type I collagen and elastin fiber assemblies establishes an important paradigm for the design in the development of novel scaffolds. In the body, both tissues and organs are organized into 3-D structures, each having specific architectures, directly dependent upon its biological function. This architecture is believed to foster cellular ingrowth and proliferation by providing appropriate channels for mass transport and spatial cellular organization, thus directing new tissue formation. The use of electrospinning technology in the arena of biomedicine has expanded the capacity of native and recombinant proteins to be fabricated into artificial extracellular matrices that more closely mimic native scaffolds. This chapter will review efforts in the development of novel structural protein materials, fabrication of nanofiber networks using electrospinning technology, and applications of subsequent structures in biomedicine.

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Electrospinning as a biomedical fabrication technology

Tissue and organ systems can be considered in general terms as a fiberreinforced composite material with associated mechanical properties largely a consequence of protein fiber networks. Electrospinning has been applied in this regard as a mechanism for generating protein nanofibers and nanofiber networks.

11.2.1 The science of electrospinning Electrospinning is a technique which relies on electrostatic forces to produce nanometer to micrometer sized fibers from polymer solutions or melts. The generation of fibers by electrospinning was first patented in 1934 by Anton Formhals for textile and polymer science (Formhals, 1934). It was not until 1977 that a revived interest in electrospinning technology emerged within the field of biomedicine for applications of wound dressings (Martin, 1977). Traditionally, engineering plastics and conducting polymers have been electrospun, but recently, with emphasis on tissue engineering and microelectronics, electrospinning protein polymers and carbon precursors has been explored. Electrospinning is essentially a drawing process utilizing electrostatic interactions in the creation of exceptionally long fibers with uniform diameters. This process is different from traditional methods of fiber formation as it is based on elongation of a viscoelastic jet of polymer solution or melt. Since elongation is achieved without contact, as with other drawing processes, this method is optimal for the development of delicate nanofibers. Electrospinning is a variant of the electrospraying technique. Both involve the use of voltage to create a jet of polymer solution. Electrospray involves the development of small droplets of low concentration solutions, a result of varicose break-up, and has been employed in areas such as mass spectrometry and ink jet printing. Alternatively, electrospinning results in the development of a small diameter fiber from highly viscous solutions. Fibers with large surface to volume ratios are produced through stretching as a consequence of the repulsion of surface charge and evaporation of the solvent (Li, 2004). There are three basic components to an electrospinning apparatus; a voltage supply, a spinneret connected to a syringe containing the polymer solution or melt, and a grounded collector. A high voltage is applied to the spinneret while the polymer solution is slowly being extruded. This induces evenly dispersed charges in a pendent drop at the tip of the spinneret, relaxing the fluid surface. This surface charge and the external Columbic forces from the electric field combine to form a tangential stress (Materials Processing Centre, 2001). This causes the drop to become distorted into a shape known as a

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Taylor cone. At a critical threshold value, the electric field strength will overcome that of surface tension and the polymer solution will be ejected as a charged jet from the spinneret tip. As the jet travels to the grounded collector it undergoes a stretching and whipping phenomena which substantially reduces the diameter of this fiber. It is then collected on the grounded collecting apparatus creating a randomly oriented nonwoven fiber network (Li, 2004). It was initially considered that the nanometer sized fibers created by electrospinning were the result of splitting and splaying of the jet due to the repulsion of surface charges (Doshi, 1995). However, with high-speed photography the ability to capture images of jet instabilities revealed that it was rapid oscillations within the jet that were responsible for stretching the fibers (Materials Processing Center, 2001). The jet was seen to travel only a few centimeters in a straight path thereupon it entered into a conical envelope and was continuously bent and whipped into a spiraling loop. As the perimeter of the loops increased, the diameter of the jet decreased (Reneker et al., 2000). Thinning of fiber diameters as much as three orders of magnitude has been noted (Fridrikh, et al., 2003).

11.3

Generation of nanofibers with controlled structures and morphology

Fiber morphology and diameter of fibers generated using the electrospinning technique is controlled by the experimental design and is dependent on both formulation and operation parameters. Critical formulation parameters include solvent selection and protein concentration, while operation parameters comprise applied voltage, gap distance, and flow rate. Through the modulation of these parameters, distinct fiber morphologies are observed. Many groups have investigated and optimized spinning conditions to produce desirable electrospun fibers. The formation of beaded fibers has been a concern and as a result of investigations by Reneker and others, options for eliminating beads have been characterized (Fong et al., 1999; Huang et al., 2000). Most fibers produced by optimized electrospinning conditions have circular crosssections while some exhibit a ribbon-like morphology, a physical difference dependent on solution concentration. For example, reports describing electrospinning of recombinant elastin proteins reveal solution concentration and flow rate to be most critical in controlling fiber morphology, as determined by high resolution microscopy. At low concentrations of recombinant elastin protein, Lys-25, (5 wt%) short, fragmented fibers were formed with a triangleor spindle-shaped beaded morphology. At higher concentrations (10 wt%) uniform fibers were generated with diameters ranging between 300–400 nm with little variation in morphology with the infrequent exception of fiber splitting at triangle-shaped bifurcation points. At 20 wt% a new morphological

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Applied voltage (b) Fiber diameter

Fiber diameter

Deposition distance (a)

Flow rate (c)

Protein concentration (d)

11.1 Experimental evidence summary plots of the relationship of electrospinning parameters to fiber morphology. (a) Fiber diameter as a function of deposition distance at constant applied voltage, flow rate and protein concentration, (b) Fiber diameter as a function of applied voltage at constant flow rate, depositon distance, and protein concentration, (c) fiber diameter as a function of flow rate at constant applied voltage, deposition distance, and protein concentration, (d) Fiber diameter as a function of protein concentration at constant applied voltage, deposition distance, and flow rate. (adapted from Bowlin, VCU, http://www.egr.vcu.edu/bme/faculty/bme-bowlin.html)

pattern was noted, defined by the emergence of flat or ribbon-shaped fibers, which appeared twisted during deposition (Huang et al., 2000). Controlling the diameter of circular fibers has been widely discussed in the literature. It has been determined that certain parameters are critical in influencing the diameter, including polymer concentration, electric field strength and flow rate. The following plots summarize the experimental evidence of the relationship of electrospinning parameters to fiber morphology (Fig. 11.1) (Fridrikh et al., 2003; Deitzel, 2001; Theron et al., 2004)

11.4

Generation of collagen and elastin small-diameter fibers and fiber networks

It is as integrated fiber networks that collagen and elastin constitute the fundamental structural elements of tissue. Thus, whether matrix proteins are produced synthetically or extracted from native tissue, it is likely that their

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versatility as scaffolds for tissue engineering applications will be significantly enhanced when reformulated into fiber networks.

11.4.1 Native collagen and biological function As a principle constituent of the extracellular matrix, collagen is ubiquitously present in all connective tissues. The most abundant form of collagen isolated from adult connective tissues, such as skin, tendon, and bone, is Type I collagen. Characteristically, it is composed of two α 1(I) chains and one α 2(I) chain, each slightly more than 1000 amino acids long, and organized as a triple helix and stabilized primarily by hydrogen bonds (Kadler et al., 1996). A single molecule of Type I collagen has a molecular mass of 285 kDa, a width of ~14 Å, and a length of ~3000 Å. In native connective tissues, Type I collagen molecules form fibrillar elements 20 to several hundred nanometers in diameter that are organized into protein networks of varying architecture. Functionally, collagen fiber networks act to resist high strain deformation and in the process transmit forces, dissipate energy, and prevent premature tissue mechanical failure (Silver et al., 1992). Collagen’s tensile strength, its stability in a biological environment, and its capacity to present specific ligands for cell surface receptors are properties that are in large measure dependent on the integrity of collagen’s characteristic triple-helical conformation (Kadler et al., 1996; Brodsky and Eikenberry 1985, Beck and Brodsky 1998; Brodsky and Ramshaw, 1997). For example, the energy requirement for collagen degradation by collagenase I (MMP-1) is reduced by a factor of two if native fibrillar (multimeric) collagen is in a monomeric form, but by a factor of ten if the triple helix is denatured (Mecham et al., 1997). Thus, while the self-assembly of monomeric collagen into an ordered supramolecular system is an important physiologic mechanism for fibril formation, the uniquely coiled triple helix is the dominant structural feature. This structure dictates collagen stability and defines its mechanical properties. Consequently, the generation of a robust load-bearing fiber network with appropriate mechanical integrity and biological function mandates maximal preservation of the collagen triple helix during the process of fiber formation. Collagen as a biomaterial Collagen is a biodegradable, biocompatible, and non-immunogenic structural protein, which makes it a suitable component for a variety of biomedical applications. As a biomaterial, collagen has been predominantly used after processing into a dry powder or slurry, a hydrogel after solution phase crosslinking, or as a porous matrix with or without the addition of other components after freeze-drying (Silver and Gary, 1997). For example, collagen has been

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used in cosmetic and urological surgery as an injectable compound for tissue augmentation (Fagien, 2000); in orthopedic surgery as an implantable matrix to promote bone growth (Parodi et al., 1997; Rao, 1995; Stol et al., 1993); and in plastic and general surgery as a topical agent for the treatment of both chronic non-healing wounds and burn injuries or as a template for tissue regeneration (Purna and Babu, 2000; Ellis and Yannis, 1996). However, it is as a native protein network that the versatility of collagen as scaffolding material could have the most profound impact in the area of tissue engineering. Notably, collagen fiber networks constitute the principle structural elements of a variety of acellular bioprosthetic tissue substitutes, such as porcine heart valves and bovine artery heterografts, as well as other tissue derived matrices, including porcine subintestinal submucosa and bovine pericardium. Electrospun collagen nanofibers The production of collagen fibers has been reported, and has traditionally relied upon wet spinning processes that involve the extrusion of a protein solution through a spinneret into an acid-salt coagulating bath, which usually contains aqueous ammonium sulfate, acetic acid, isopropanol, or acetone. Further treatments in ethanol and acetone solutions are often required for fiber dehydration. For example, Hirano et al. (2000) described the production of chitosan-collagen fibers (d > 30 µm) produced by wet spinning from an aqueous acetic acid-methanol solution into an ammonia solution containing 40–43% ammonium sulfate. Likewise, Fofonoff and Bell (1999) have reported a method for forming collagen fibers (d > 100 µm) by wet spinning from an aqueous acetic acid solution into a heated coagulating bath containing alkaline alginic or boric acid. The collagen fiber is formed by polymerization when the acid in the collagen is neutralized upon contact with the neutralizing solution and the fibers are subsequently dehydrated in acetone and ethanol baths. An additional example is provided by Furukawa et al. (1994) in which solubilized collagen is spun into a coagulating bath containing an inorganic salt, such as sodium, aluminum, or ammonium sulfate. Nonetheless, limitations of these approaches are recognized, including: (a) the use of conditions which likely induce significant conformational changes in native protein structure, including protein denaturation; (b) the generation of fibers that range from tens to hundreds of microns in diameter and are much larger than those observed in native tissues (Merrilees et al., 1987; Buck, 1987; BreitenderGeleff et al., 1990); and (c) a reliance on biologically toxic solvent systems. Although research in the area of wet spinning collagen has advanced and significant improvements have been achieved, an alternate approach for submicron collagen fiber formation, electrospinning, has recently been investigated (Stitzel et al., 2006; Li et al., 2005; Buttafoco et al., 2006; Zhong et al., 2005; Zhong et al., 2006; Matthews et al., 2002, Huang et al.,

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2001). The architecture generated from this process is similar to that found in most native extracellular matrices, thus underscoring the electrospinning technique for design of novel scaffolds. The first report of electrospun collagen fibers employed a weak acid solution to electrospin Type I collagen-polyethylene oxide (PEO) blends at ambient temperature and pressure. High resolution microscopy was employed to resolve the influence of critical electrospinning parameters, specifically, solution viscosity, conductivity, and flow rate on subsequent fiber ultrastructure and size. A variety of fiber microstructures were observed: beaded, round and ribbon-like filaments. Ultimately, fibers of uniform morphology and ultrastructure, with average diameters of 100–150 nm, were generated. Significantly, this procedure outlined a non-toxic and non-denaturing approach for the generation of collagen containing nanofibers and nonwoven fiber networks (Fig. 11.2) (Huang et al., 2001). Similarly, other approaches have investigated various collagen sources and isotypes in the production of collagen nanofibers. Typically, acid soluble Type I collagen from rat tail tendons or calf skin have been utilized. Type I and Type III collagen from human placenta have also been investigated (Matthews et al., 2002). Results indicate that identity and source of collagen are significant to the morphological, mechanical, and biological properties of the electrospun collagen networks. Additionally, solvents such as HFIP (1,1,1,3,3,3 hexafluoro-2-propanol) have been used for electrospinning of collagen. While some investigators have claimed preservation of native collagen structure, studies in our own laboratory demonstrate complete loss of triple helical structure when examined by circular dichroism spectroscopy, differential scanning calorimetry, or x-ray diffraction (Buttafoco et al., 2006; Rho et al., 2006; Huang et al., 2001).

11.5

Biological role of elastin

Native elastin is a highly insoluble matrix protein which functions to provide extensibility and resilience to most tissues of the body. Elastin networks are responsible for maximizing the durability of tissues that are loaded by repetitive forces by minimizing the conversion of mechanical energy to heat which ultimately results in tissue damage (Lillie and Gosline, 2002). In addition to the structural role, elastin creates an environment, that promotes proper cell function and modulates cellular attachment, growth, and responses to mechanical stimuli. Elastin fibers appear to exist as two morphologically different components; a highly isotropic amorphous elastin constituent within an organized microfibrilar scaffold (Alberts et al., 2002). Understanding of the mechanism of fiber assembly in native elastin is limited; however, it appears to take place in proximity to the cell membrane where microfibrils emerge as fiber

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11.2 SEM micrographs of PEO-collagen blended fibers spun from 2 wt% acid solution (34 mM NaCl) at a flow rate of 100 µl min–1 and at different collagen–PEO weight ratios: (a) 30 : 1, 50 000x magnification, (b) 10 : 1, 50 000x magnification (c) 5 : 1, 50 000x magnification, (d) 2 : 1, 50 000x magnification, (e) 1 : 1, 20 000x magnification, (f) 1 : 2, 50 000x magnification. Fibers of uniform morphology and ultrastructure, with average diameters of 100–150 nm, were generated (adapted from Huang et al., 2001)

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bundles. Amorphous elastin is synthesized by smooth muscle cells as a soluble monomer, the 72 kDa precursor tropoelastin, and is secreted within each fiber bundle. Similarly to natural rubber, it is organized into insoluble networks reminiscent through enzymatic crosslinking via oxidation by lysyl oxidase (Garrett and Grisham, 1999). The distinctive composition of tropoelastin affords unique physical properties of this structural protein. Tropoelastin is rich in glycine (33%), proline (10– 13%), and other hydrophobic residues (44%) rendering elastin an extremely hydrophobic protein (Rucker and Dubick, 1984). Tropoelastin contains distinct crosslinking and hydrophobic domains. Crosslinking domains are alanine rich, containing pairs of lysine residues thereby facilitating intermolecular crosslinking. Alternatively, the hydrophobic domains within tropoelastin are composed of three-quarters valine, glycine, proline and alanine. Investigations have elucidated that the precise sequence and size of this region are not critical for appropriate function; however, the total size of the protein polymer, 750–800 residues, is highly conserved among species (Rosenbloom et al., 1993).

11.5.1 Elastin as a biomaterial A failure of current acellular bioprostheses is their inability to exhibit mechanical properties that match those of native tissues, primarily a result of the loss or degradation of the elastin protein networks, thereby reinforcing the importance of elastin fiber networks is bioprosthetic design. Isolated elastin matrices from acellular allo- and xenogenic tissues have been investigated as scaffolding materials with these studies confirming that native protein fiber networks can be used to fabricate an artificial scaffold. However, these scaffolds often require the addition of structural proteins or must be seeded with cells to demonstrate proper biochemical and biomechanical function (Berglund et al., 2004; Lu et al., 2004). Despite successes, recognized drawbacks, including tissue heterogeneity, incomplete cell extraction, the generation of ill-defined chemical crosslinks, progressive biodegradation, and the potential risk of viral transmission from animal tissue, continue to dampen enthusiasm for this approach. As a promising alternative in the generation of biomimetic scaffolds, soluble elastin, derived either as fragmented elastin, in the form of alpha- or kappa-elastin, or as the natural monomer tropoelastin (Li et al., 2005), have been successfully electrospun. Additionally, through genetic engineering of synthetic polypeptides, novel elastin proteins have been created for such applications. Utilizing these strategies affords the ability to tailor matrix composition and content, fiber size and architecture, or other features that may influence 3-D hierarchical tissue structure, thus enabling the ability to design a scaffold with precisely defined mechanical and biological properties.

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11.5.2 Recombinant elastin technologies It has been postulated that the generation of protein polymers that mimic native structural proteins and the assembly of these recombinant proteins either alone or in combination with naturally occurring matrix proteins provides an opportunity to optimize the mechanical properties of artificial tissues. In this way, recombinant technologies have been pursued in the generation of elastin-mimetic protein polymers. Through the structural characterization of the hydrophobic domains, the ability to base synthetic protein polymers on native elastin sequences is feasible. The pioneering work of Urry elucidated the elastomeric pentapeptide repeat VPGVG, from human elastin, which now serves as the fundamental sequence extensively investigated by both chemical methodologies and recombinant technology (Urry, 1997; Urry, 1998). VPGVG is a common repeat unit within the hydrophobic domain of human elastin and is responsible for resultant elastic properties. Additionally, this domain is responsible for facilitating fiber formation through coacervation phenomena, behaviors consistent with native elastin. Spectroscopic analysis has revealed that native elastin, and likewise, protein polymers containing this repeat, exhibit β-turns and helical β-spiral conformations and display an inverse temperature transition defined by the generation of a more ordered system upon increasing temperature. This loss of entropy is a consequence of protein folding into β-spiral conformation and the subsequent reorientation of water from the elastin chain (Chang and Urry, 1988). Studies have elucidated the amino acid in the fourth (X) position (VPGXG) modulates the coacervation temperature with more polar amino acids increasing transition temperature (Urry et al., 1991; Urry et al., 1992; van Hest and Tirrell, 2001). Preservation of the glycine and proline residues maintain the structure and function of elastin analogs (van Hest and Tirrell, 2001). This discovery has led to the generation of recombinant elastin analogs designed for biomedical applications. For instance, this technology has been employed in the design of amphiphilic elastin protein polymers consisting of hydrophobic and hydrophilic domains. Through precise sequence design and control of processing conditions, these elastin analogs exhibit a wide range of properties advantageous for biomedical applications, as micelles, physically crosslinked hydrogels, or nanofiber networks (Wright and Conticello, 2002; Wright et al., 2002; Wu et al., 2005; Nagapudi et al., 2005; Huang, 2000). Additionally, groups have incorporated cell binding domains, RGD or REDV, into elastin sequences to functionalize elastin matrix components for endothelial cell attachment (Panitch et al., 1999; Welsh and Tirrell, 2000). Genetic engineering strategies afford the ability to modulate macroscopic properties on the molecular level. Therefore the potential exists to generate synthetic polypeptides that mimic native proteins. In this regard, there is an inherent opportunity to precisely engineer recombinant sequences to targeted design criteria such as tensile strength,

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elastic modulus, viscoelasticity, and in vivo stability, as well as the optimization of a desired host response.

11.5.3 Generation of elastin and elastin-mimetic small diameter fibers and fiber networks As material for tissue engineering applications, elastin is intended to provide both mechanical support and potentially act as a scaffold for cellular repopulation. As such, it is likely when reformulated into fiber networks that the versatility of elastin as a scaffolding material will be significantly improved. In this regard, electrospinning has been investigated as a mechanism for generating fibers with diameters < 1 µm. When proteins are reformulated as fiber systems desired mechanical and biological properties can be achieved for biomedical applications. For instance, flexibility of a fibrous system can be controlled by either a decrease in fiber diameter or an increase in fiber number (Ottani et al., 2001). Thus, reformulating elastin proteins into fiber networks provides an additional level of control over the properties of the artificial matrix designed. Specifically, studies have indicated electrospun fabrics composed of small diameter fibers (<1 µm) to have decreased porosity, increased fiber density, increased mechanical strength, as well as an optimized biological environment for promoting cell adhesion as compared to larger diameter fibers (7 µm) (Li et al., 2005; Kwon et al., 2005). Previous reports have demonstrated the feasibility of electrospinning soluble elastin as a single component system as well as in collagen-elastin blends. Concentrated solutions of soluble elastin produced fibers with average diameters from ranging from 100 nm to 3 µm, with fiber diameter being highly dependent upon solvent systems utilized (Buttafoco et al., 2005; Buttafoco et al., 2006; Li et al., 2005; Stitzel et al., 2006). Additionally, collagen-elastin blends produced fibers with diameters ranging from nanometers to micrometers (Boland et al., 2004; Buttafoco et al., 2006). Subsequent work has endeavored to improve the mechanical properties, specifically the compliance and mechanical strength of these elastin based matrices, through the addition of synthetic polymeric materials such as PLGA, poly (D,L-lactide-co-glycolide). Notably, compliance testing of collagen-elastin-PLGA electrospun scaffolds demonstrated behavior consistent with in vivo mechanical behavior of bovine arteries. Specifically, controlling the ratio of collagen, elastin, and PLGA facilitated the modulation of electrospinning characteristics as well as the strength and stability of the electrospun scaffold, as a burst pressure of nearly 12 times normal systolic pressure was observed (Stitzel et al., 2006). Recently, attempts to electrospin human tropoelastin have been undertaken. Interestingly, electrospinning of these materials produced ribbon-shaped fibers, several microns in width, which appeared to retain a unique periodicity,

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controlled by modulating the solution flow rate. Morphology of these fibers is reminiscent of the topology of elastin in native tissues; for instance, the architecture of the elastic lamina within blood vessels (Li et al., 2005) Notwithstanding reports with promising results, important limitations remain associated with native elastin electrospinning strategies. In consideration of these, significant investigations in the design of recombinant elastin proteins, microbial expression of these proteins, and reformulation into protein fiber networks by means of electrospinning have been explored. The first report of fiber formation from an elastin-like analog was described utilizing a 81 kDa recombinant elastin peptide polymer, Lys-25, comprised of the repeat sequence (Val-Pro-Gly-Val-Gly)4(Val-Pro-Gly-Lys-Gly). Electrospinning of a 15-wt% solution afforded a fabric with a unimodal distribution of fiber diameters (Fig. 11.3). Moreover, in the absence of high rotational or translational speeds on a collecting mandrel, fiber orientation was random. Interestingly, pulsed field gradient NMR spectroscopy was utilized to access network porosity and pore size distribution in 3-D fabrics and revealed Lys-25

(a)

(b)

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11.3 First reported fiber formation from an elastin-like analog utilizing a 81 kDa recombinant elastin peptide polymer. SEM micrographs of elastin-mimetic peptide fibers spun from 15 wt % solution at 50 (a), 100 (b), 150 (c), and 200 µl min–1 (d) flow rate. Electrospinning of a 15-wt% solution afforded a fabric with a unimodal distribution of fiber width (adapted from Huang et al., 2000).

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electrospun matrices as physiologically ideal for cellular seeding in that regard (Fig. 11.4) (Huang, 2000; Nagapudi et al., 2002). While a number of elastin-mimetic protein polymers can be fabricated as fibers and covalently crosslinked, a family of recombinant elastin triblock copolymers containing chemically distinct midblocks has been investigated in the generation of virtually crosslinked fibers and fiber networks. Specifically, these studies demonstrate that self-assembling triblock elastin analogs have the capacity to form stable fibers, but without a requirement for chemical crosslinking (Nagapudi et al., 2005). As an example, electrospinning was used to produce tubular conduits from a triblock copolymer. Mechanical properties were assessed in PBS at 37°C. Hydrated samples displayed an elastic modulus of 0.29 ± 0.03 MPa and a strain to failure of 151 ± 29%, values comparable to

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11.4 (a) Diffusion NMR data for a fabric sample made from an elastin-like analog as a function of diffusion time. (b) Fit shown for a Gaussian distribution of pore sizes on the 1000 ms diffusion data. Average pore diameter was found to be 78 µm (adapted from Nagapudi et al., 2002).

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the elastin component of the arterial wall (Young’s modulus ~ 0.3 MPa) (Urry, 1984; Niklason et al., 1999).

11.6

Generation of crosslinked fibers and fiber networks

It is generally understood that crosslinking is necessary for the maintenance of biostability in engineered collagen and elastin scaffolds. In native tissues, lysyl oxidase, a specific amine oxidase, catalyzes the formation of aldehyde cross-link intermediates in the solid state providing intermolecular crosslinking within collagen and elastin networks. Presumably, the incorporation of various degrees of crosslinking can be used to further tailor and control the material properties of the scaffold to specific applications.

11.6.1 Crosslinking collagen networks Glutaraldehyde is the traditional crosslinking agent for bioartificial devices. While it has been shown to enhance stability and further reduce immunogenicity, it does exhibit potential cytotoxicity and late induction of collagen calcification. Therefore, alternative crosslinking approaches have been reported, including chemical (carboiimide, diisocyanates and polyepoxy compounds) and solid state photocrosslinking. For example, Chaikof and colleagues described the derivitization of Type I collagen with methacrylate groups (Brinkman et al., 2003; Nagapudi et al., 2002) and cinnamate groups for photocrosslinking (Dong et al., 2005) while preserving collagen’s triple helical structure.

11.6.2 Crosslinking elastin networks Crosslinking of native elastin, as well as synthetic elastin-mimetic protein polymers, has most often been investigated using solution phase systems; either gamma irradiation (Lee J et al., 2001b; Lee et al., 2001c), chemical (Lee et al., 2001a; Nowatzki and Tirrell, 2004; Trabbic-Carlson et al., 2003; Nagapudi et al., 2002), or enzymatic based approaches (Kagan et al., 1980), as well as solid state photocrosslinking (Nagapudi et al., 2002). Nevertheless, for most biomedical applications of synthetic scaffolds, vapor phase glutaraldehyde crosslinking is currently the system utilized, with successful in vitro and in vivo biocompatibility results. Directed efforts in alternative chemical crosslinking strategies, in which no additional chemical entities are introduced and stable amide linkages are formed, have also been investigated. For example, mixed solutions of EDC (N-(3-dimethylaminopropyl)-N’ethylcarboiimide hydrochloride) and NHS (N-hydroxysuccinimide) were employed to crosslink electrospun collagen-elastin scaffolds with no apparent

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alteration to fiber morphology as determined by microscopy analysis. Scaffold stability and biocompatibility was assessed in vitro through cell seeding of smooth muscle cells with promising results (Buttafoco et al., 2006). Alternatively, incorporation of reactive lysine residues into recombinant elastin design provides the ε-amino moiety of lysine for crosslinking using a variety of approaches. This strategy affords the ability to achieve precise control over the nature and degree of crosslinking, and facilitates spatial and temporal control over the reaction process. Specific investigations into reactive group spacing as well as crosslinking strategies on the modulation of important biological behaviors of elastin analogs has been conducted with the general conclusion that the placement of well defined crosslinks enhance the biostability of elastin and improve biologically relevant properties. For example, methacrylate groups were incorporated into the protein polymer backbone in order to facilitate site-specific solid-state photocrosslinking using either UV or visible light activated photoinitiators (Nagapudi et al., 2002). Mechanical analysis confirmed superior biologically relevant behaviors, comparable to the elastic behavior of native elastin. Significantly, in response to the deleterious effects of chemical crosslinking reagents, a new class of recombinant proteins have been investigated: selfassembling triblock elastin copolymers, which have the capacity to form stable fibers, but without a requirement for chemical crosslinking (Nagapudi et al., 2005; Wright and Conticello, 2002; Wright et al., 2002). Due to the nature of the copolymer design, they form physically or virtually crosslinked systems. Notably, nanofiber formation was influenced by solvent conditions, with nanofibers in the diameter range of 100–400 nm generated utilizing a solvent which identically solvates both blocks of the copolymer, while nanofibers with diameters ranging from 800 nm–3 µm were generated from an aqueous solution which preferentially solvates only the midblock. Subsequent mechanical evaluation indicated that modulation of elastic and plastic behavior of the triblock protein was directly dependent on the solvent systems used in fabrication. In particular, mechanical evaluation of the triblock elastin protein fibers under physiological conditions revealed elastic modulus and ultimate tensile strength comparable to that of native tissues such as the elastin component of the arterial wall (Young’s modulus ~ 0.3 MPa) (Urry, 1984; Niklason et al., 1999).

11.7

Multicomponent electrospun assemblies

It is likely that the molecular structure and supramolecular organization of collagen and elastin fiber assemblies within native tissues establishes an important paradigm for the design of a biomimetic scaffold. Therefore, the hierarchal assembly of these protein fiber networks is essential for generating constructs with both enhanced biostability and blood contactability, as well

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as mechanical properties that closely match those of the native tissue. As such, the potential of electrospinning multicomponent systems in the design of biomimetic scaffolds is substantial since it affords an additional level of control in vitro to recreate native architectures. In this regard, two electrospinning techniques have been employed to create nanofibrous bicomposites: multilayering electrospinning and multicomponent (mixing) electrospinning (Kidoaki et al., 2005). Multilayering electrospinning requires sequential deposition of two unique materials onto the same collector, such that a hierarchal ordered structure is obtained. This method has been employed to fabricate constructs composed of synthetic polymers lined with biocompatible native proteins such as collagen, fibrin and laminin. As an alternate technique, mixing electrospinning, in which two unique materials are simultaneously electrospun from different syringes, ultimately yields a fiber network of mixed identity. This method is of particular interest in creating tubular constructs for tissue engineering of blood vessels. For instance, a bilayered construct can be engineered with an outer layer of circumferentially oriented fibers mimicking the tunica media and an inner layer of randomly oriented fibers acting as the elastic lamina of the blood vessel (Vaz et al., 2005). This method has also been employed to modulate porosity and microvoid spaces within an electrospun scaffold through fiberleaching. Several such studies involving simultaneous spinning of PEO and other synthetic or natural polymers, followed by dissolution of the PEO component in water have been investigated in the modulation of scaffold architecture (Kidoaki et al., 2005; Huang et al., 2001). Additionally, blended solutions have been investigated, specifically collagenelastin blends, which yield a fiber matrix in which singular components are indistinguishable. These types of matrices have exhibited enhanced mechanical and biological behaviors compared to matrices composed of the individual components (Buttafoco et al., 2005; Stitzel et al., 2006). Collagenglycosaminoglycan scaffolds have also been electrospun from blended solutions with a pore structure (mean diameter of 260 nm) similar to that found in native matrices. Upon vapor phase crosslinking, these scaffolds exhibited biostability and resistance to collagenases along with increased cellular proliferation when seeded with cells (Zhong et al., 2005).

11.7.1 Electrospun nanofiber networks and the potential for the incorporation of living cells Both collagen and elastin matrix proteins provide a useful physiological starting point for the creation of a biochemical and biomechanical environment that is optimized for enhanced cell adhesion, migration, proliferation, and

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differentiation. Fabrication of these proteins into nanofiber networks provides the potential to incorporate various cell types – endothelial cells, SMC, fibroblasts, stem cells, chondrocytes, osteoblasts, human or animal cells – which have been genetically engineered to produce a protein of interest (e.g., growth factor, peptide hormone, antiangiogenic protein). In vitro investigations indicate the architecture of electrospun networks supports mass transport and spatial cellular organization. Specifically, as a consequence of the fiber nano-dimensions, the surface-to-volume ratio, and the unique three-dimensional architecture within these electrospun matrices, movement of signaling molecules, nutrients, and metabolic waste is likely enhanced and cell-cell and cell-matrix interactions are facilitated (Smith and Ma, 2004). Specifically, collagen and elastin scaffolds have been shown to promote cell adhesion, migration and proliferation of a variety of cell types in vitro, with promising results concerning biocompatibility and biostability.

11.8

Future trends

There is a long felt need for durable materials for medical and veterinary use in organ and tissue substitutes. Such materials must be compatible with human and animal physiologies such that thromboses, inflammation and other harmful physiological reactions are not induced. Furthermore, durable and biologically compatible materials which do not require lengthy preconditioning periods prior to implantation are required. Over the past decade, considerable effort has been directed towards developing scaffolds using both synthetic and natural polymers. As the majority of human tissues and organs originate from hierarchically organized fibrous structures, electrospinning fiber networks is of particular interest in the development of these unique matrices. Importantly, scaffold geometry can be modulated for a variety of tissue engineering applications, simply by the shape of the collector, such that seamless and complex scaffold geometries can be fabricated. Additionally, from a commercial distribution perspective, the electrospinning technique is a rapid and efficient technology which can easily be conducted utilizing a sterile technique to generate a material which will likely have a long shelf life. The many diverse areas of research in which electrospinning is being employed within biomedicine underscores the range of applications for which this technology can be utilized. These matrices not only mimic native ECM architecture, which ultimately reproduces native mechanical and biological performance, but also provides the opportunity to further tailor biological responses. Additionally, the electrospinning technique provides enormous flexibility to tissue engineering of biocompatible matrices for a variety of applications. These matrices have potential in drug delivery (Dong et al., 2005), vascular bioengineering of blood vessels and heart valves (Stitzel

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et al., 2001; Boland et al., 2004), hard and soft tissue reconstruction, load bearing prosthetic materials, materials to facilitate wound closing and/or healing and stem cell delivery (Keen et al., 2004).

11.9

References

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Zhong S, Teo W E, Zhu X, Beuerman R, Ramakrishna S and Yung L Y (2005), Formation of collagen-glycosaminoglycan blended nanofibrous scaffolds and their biological properties, Biomacromolecules, 6, 2998–3004. Zhong S, Teo W E, Zhu X, Beuerman R W, Ramakrishna S and Yung L Y (2006), An aligned nanofibrous collagen scaffold by electrospinning and its effects on in vitro fibroblast culture, J Biomed Mater Res A, 79, 456–463.

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12 Starch-polycaprolactone based scaffolds in bone and cartilage tissue engineering approaches M. E. G O M E S, J. T. O L I V E I R A, M. T. R O D R I G U E S, M. I. S A N T O S, K. T U Z L A K O G L U, C. A. V I E G A S, I. R. D I A S and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

12.1

Introduction

Most of the tissue engineering strategies developed for the creation of hard tissue (such as bone and cartilage) substitutes relies on the use of a temporary three dimensional (3-D) scaffold material in which cells are seeded and in vitro cultured prior to implantation. In this type of strategy, the formation of new tissue is deeply influenced by the 3D environment provided by the scaffolds, namely their composition, porous architecture and, of course, their biological response to implanted cells and/or surrounding tissues. In order to meet all the necessary requirements, scaffold materials must be fabricated from polymers with adequate properties, but many of the scaffold’s features are also dictated by the processing methodology used to fabricate them. Biocompatibility is the first obvious demand, but an ideal tissue engineering scaffold should also exhibit appropriate mechanical properties (Freed and Vunjak-Novakovic, 1998; Middleton and Tipton, 2000; Vacanti and Bonassar, 1999; Kim and Mooney, 1998; Chapekar, 2000; Thomson et al., 1995) and a suitable degradation rate (Thompson et al., 1995; Middleton and Tipton, 2000; Kim and Mooney, 1998; Chapekar, 2000; Thomson et al., 1995; Hutmacher, 2000). Furthermore, the scaffold must possess adequate porosity, interconnectivity and permeability to allow the ingress of cells and nutrients (Hutmacher, 2000; Thompson et al., 1995b; Kim and Mooney, 1998; Chapekar, 2000) as well as the appropriate surface chemistry for enhanced cell attachment and proliferation (Langer, 1999; Kim and Mooney, 1998; Chapekar, 2000; Freed and Vunjak-Novakovic, 1998). The recent development of several different processing methodologies and polymers enables the creation of scaffolds with a wide range of structures and geometries, as well as mechanical properties and degradation profiles, among other characteristics. Fiber bonding methodologies produce fiber mesh scaffolds that consist of individual fibers either woven or knitted into 3D patterns of variable pore sizes (Thompson et al., 1995a; Langer, 1999; Lu and Mikos, 1996; Maquet and Jerome, 1997; Thompson et al., 1997). Fiber 337 © 2008, Woodhead Publishing Limited

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meshes usually exhibit a large surface area for cell attachment, which also enables a rapid diffusion of nutrients enhancing cell survival and growth (Thompson et al., 1995a; Langer, 1999; Lu and Mikos, 1996; Maquet and Jerome, 1997; Thompson et al., 1997). This, of course, results from a high interconnectivity among pores, that contrasts with the difficulty in controlling accurately the porosity (Thompson et al., 1995a; Lu and Mikos, 1996; Maquet and Jerome, 1997; Thompson et al., 1997). These features make this a very attractive method for the production of scaffolds for tissue engineering. Fiber bonding methods include a great variety of processing methods that involve the knitting or physical bonding (by means of casting or compression procedures) of fibers prefabricated by wet or dry spinning from polymeric solutions or by melt spinning. Fiber meshes may also be obtained in single step methods such as electrospinning. Several studies demonstrate that scaffolds obtained by fiber bonding processes have adequate structure for tissue engineering strategies using bioreactor cultures, probably because they provide highly interconnected porosity which creates hydrodynamic micro-environments with minimal diffusion constrains that closely resemble natural in vivo interstitial fluid conditions, achieving large and well organized cell communities. In contrast most of the pores obtained with other methodologies exhibit lower interconnectivity, which is very likely to generate complex fluid flow pathways thought the scaffolds, and that does not allow the distribution of cells throughout the whole construct. In this chapter we described the use of SPCL (starch+ε-polycaprolactone, 30/70%) fiber meshes, in bone and cartilage tissue engineering strategies based on the in vitro culturing of the scaffolds with different types of cells under static and flow perfusion conditions, previous to implantation. The scaffolds based on SPCL were prepared from fibers obtained by melt-spinning by a fiber bonding process in order to produce fiber meshes with different porosities. The porosity of the scaffolds was characterized by microcomputerized tomography (µCT) and scanning electron microscopy (SEM). The scaffolds’ degradation behavior was assessed in solutions containing hydrolytic enzymes (α-amylase and lipase) in physiological concentrations, in order to simulate in vivo conditions.

12.2

Starch+ ␧-polycaprolactone (SPCL) fiber meshes

Fiber mesh scaffolds based on SPCL (a 30/70 wt% blend of starch with poly (α-caprolactone)) were prepared by a fiber bonding process consisting of cutting and sintering melt-spun fibers with a diameter of about 180 µm. The different porosity of the fiber meshes was obtained using different amounts (by weight) of fibers. The SPCL scaffolds obtained by this method exhibit a

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typical fiber-mesh structure, with highly interconnected pores and a porosity of approximately 50% or 75%, as determined by µCT analysis. All samples were cut into disks of approximately 6 to 8 mm diameter and 1.5 to 2 mm height and were sterilized using ethylene oxide, before being used in the cell culture studies. PCL is a biodegradable aliphatic polyester with important applications in the biomedical area (Chastain et al., 2006; De Jong et al., 2005; Williams et al., 2005) and several studies have been performed on the degradation behavior of this biomaterial (Catiker et al., 2000; Chawla and Amiji, 2002; Li, 1999, Liu et al., 2000; Xiao et al., 2003; Tsuji, 2005; Pena et al., 2006). PCL hydrolysis may be catalysed by lipase enzymes (Chawla and Amiji, 2002; Liu et al., 2000; Tsuji et al., 2006; Pei et al., 2002). The natural function of lipases is the hydrolysis of triglycerides to partial glycerides and fatty acids. Serum lipase is mainly derived from the pancreatic acinar cells but other sources of lipase in the human body are the digestive tract, adipose tissue, lung, milk and leukocytes (Tietz and Shuey, 1993). The serum lipase concentration in healthy adults is in the range of 30-190 U/L (Tietz and Shuey, 1993). On the other hand, α-amylase enzyme is able to catalyse the hydrolysis of α-1,4-glycosidic linkages of starch, reducing the molecular size of starch and producing maltose and dextrins. In humans, the enzyme occurs in a variety of tissues, but the highest concentrations are found in the pancreas and in salivary glands. Low amylase activities are normally detected in the serum (46-244 U/L) (Junge et al., 1989) of healthy subjects. Taking into account the catalytic activities of lipase and α-amylase enzymes, the degradation behavior of SPCL scaffolds was studied using a PBS solution and PBS containing one or both enzymes.

12.3

SPCL-based scaffold architecture, stem cell proliferation and differentiation

As previously mentioned, adequate porosity and surface area are widely recognized (Leong et al., 2003; Ma and Choi, 2001; Murphy et al., 2002; Shea et al., 2000) as important parameters in the design of scaffolds for tissue engineering. Other architectural features such as pore morphology and interconnectivity between pores of the scaffolding materials are also suggested to be important for cell seeding, migration, proliferation, mass transport, gene expression, and new 3D tissue formation. (Leong et al., 2003; Ma and Choi, 2001; Wake et al., 1994). In traditional static cell culture systems, cells within a scaffold construct will receive nutrients only by diffusion from the surrounding media (Wang et al., 2003; Botchwey et al., 2001). Thus, with static culture, high cell densities on the scaffolds’ exteriors often deplete nutrient supplies before these nutrients can diffuse to the scaffold interior (Wang et al., 2003). These

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diffusion limitations in static culture become more pronounced with the decreasing of scaffold porosity. Cells in the interior of statically cultured scaffolds may ultimately become necrotic due to a lack of nutrients and an excess of metabolic waste products (Wang et al., 2003; Shea et al., 2000). In addition to these transport restrictions, static culture systems also fail to provide any mechanical stimulation to cell seeded scaffolds. It is well known that bone cells are sensitive to mechanical stimulation (Butler et al., 2000; Cowin, 1998), constant in their natural environment. The absence of mechanical stresses may therefore hinder in vitro cell development (Butler et al., 2000). These findings have motivated the development of enhanced culture systems, such as the flow perfusion bioreactor (Wang et al., 2003; Bancroft et al., 2002, 2003; Temenoff and Mikos, 2000), which simultaneously provides sufficient transport of nutrients and wastes and a continuous mechanical stimulation of cells (Wang et al., 2003; Temenoff and Mikos, 2000; Freed and Vunjak-Novakovic, 2000) creating dynamic culture environments that support in vitro formation of 3-D bone-like tissue. Previous studies (Gomes et al., 2006b) investigated the influence of the porosity of fiber mesh scaffolds obtained from a blend of starch and poly(ε-caprolactone) on the proliferation and osteogenic differentiation of marrow stromal cells cultured under static and flow perfusion conditions. For this purpose, SPCL scaffolds were fabricated into mesh structures with two different porosities, namely 50% and 75%. These scaffolds were then seeded with marrow stromal cells harvested from Wistar rats and cultured in a flow perfusion bioreactor or in 6-well plates for up to 15 days. It was also demonstrated that increased scaffold porosity significantly enhances the proliferation of marrow stromal cells both cultured under static and flow perfusion conditions and influences the sequential development of the seeded cells. Furthermore, the flow perfusion induces de novo tissue modeling with the formation of pore-like structures in the scaffolds with higher porosity (75%), demonstrating that this structural aspect of scaffolding materials, in combination with the culture environment, determines, to a great extent, the structure and possibly the functionality of bone-like tissue substitutes formed in vitro. In fact, biodegradable starch-based fiber mesh scaffolds in conjunction with fluid flow bioreactor culture enable the creation of culture environments with minimal diffusional constraints and the ability to provide mechanical stimulation to seeded marrow stromal cells, leading to an enhancement of their differentiation towards the development of a bone-like extracellular matrix and its mineralization, forming a carbonated apatite mineral similar to the major mineral component of bone. In another study (Gomes et al., 2006a), we have hypothesized that, since flow perfusion bioreactors create culture environments with minimal diffusion constraints, providing cells with mechanical stimulation may closely resemble in vivo conditions for bone formation. Therefore, these culturing systems, in conjunction with an appropriate scaffold and cell type, may provide significant

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insight towards the development of in vitro tissue engineering models leading to improved strategies for the construction of bone tissue substitutes. Therefore, we have investigated the in vitro expression of several bone growth factors that are usually associated with in vivo bone formation by culturing rat bone marrow stromal cells seeded onto starch-based biodegradable fiber meshes in a flow perfusion bioreactor. Immunohistochemical analysis of cell-scaffold constructs cultured in the flow perfusion bioreactor for different time periods consistently showed the presence of positively stained regions for all the growth factors examined (namely for BMP-2, FGF-2, VEGF and TGF-β1) except for PDGF-A. The growth factor expression is enhanced with an increasing flow rate due to the enhanced differentiation induced by mechanical stimulation of the cells. A trend for increased immunohistochemically stained area over culturing time was also observed. These results provide evidence that growth factors can be delivered into a scaffold via co-transplantation of cells that can naturally release them when cultured in stimulating conditions and thus accelerate the healing and/or neotissue development upon implantation of the construct. In this sense, flow perfusion augments the functionality of scaffold/cell constructs grown in vitro as it combines both biological and mechanical factors to enhance cell differentiation and cell organization within the construct (Gomes et al., 2006a). This study also shows that flow perfusion bioreactor culture of marrow stromal cells combined with biodegradable starch-based fiber meshes may constitute a useful model for in vitro studies on the biological mechanisms associated with bone formation and regeneration. In fact, the true biological environment of a bone cell is derived from a dynamic interaction between responsively active cells experiencing mechanical forces and a continuously changing 3D matrix architecture, which can be simulated, obviously to a limited extent, in this type of bioreactor (Gomes et al., 2006a). Therefore, the additional influence of the biodegradable scaffold used in this system cannot be excluded. In this case, it seems that SPCL fiber meshes support the expression of the different growth factors by rat bone marrow stromal cells cultured within these scaffolds, providing further evidence of their suitability for bone tissue engineering applications.

12.4

In vivo functionality of SPCL fiber-mesh scaffolds

The ability of a tissue engineering approach to promote the formation of a new functional tissue and thus bone defect regeneration must be ultimately assessed in vivo. Cell based strategies sustained by a support material have been applied to generate ectopic or ortotopic bone (Meinel et al., 2006; Kruyt et al., 2007). Although the latter presents a major potential for skeletal

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regeneration procedures, most of the in vivo studies are conducted in nude mice or rats and/or using an ectopic approach (Livingston et al., 2002; Mendes et al., 2003; Trojani et al., 2006; Kruyt et al., 2007; Mauney et al., 2005; Mastrogiacomo et al., 2007). Autologous approaches have also been considered in recent studies (Zhu et al., 2004; Kruyt et al., 2004; Niederauer et al., 2000) avoiding immune complex problems that interfere with the regenerative process as well as with the patient follow up. The in vivo functionality of osteogenic tissue engineered constructs obtained by in vitro culture of marrow stromal cells onto starch-polycaprolactone (SPCL) scaffolds at different stages of development, was assessed using an autologous approach in a goat model (Rodrigues et al., 2008). For this purpose, goat marrow stromal cells (GBMCs) harvested from the iliac crests of adult goats were expanded (using autologous serum) for 2 to 3 weeks before seeding them onto SPCL scaffolds (d = 6 mm/h = 2 mm). The cell-seeded scaffolds were then in-vitro cultured for 1 or 7 days prior to implantation using osteogenic medium. In vitro cell proliferation and differentiation were analyzed for the same culturing periods by DNA quantification and ALP activity respectively. Non-critical size defects were drilled in two femurs of four goats from which the cells were harvested: per femur, two defects were left empty, two defects were filled with the SPCL alone (controls), and the remaining were filled with GBMCs seeded and cultured onto SPCL for one or seven days, respectively. Xylenol orange, calcein green and tetracycline markers were injected subcutaneously two, four and six weeks after implantation, respectively, to assess different bone neoformation stages. Six weeks after implantation, the animals were euthanized, and femurs cut into single defect-sections and observed for fluorescence detection or Lévai Laczkó staining. Implanted constructs showed no significant inflammatory response during the regeneration process. Neobone was observed in all defects but observations suggest a higher bone growth in the defects filled with cells-SPCL constructs when compared to empty- or SPCL alone- defects. Bone growth may be enhanced by the presence of cell-scaffold constructs and the in vitro culturing time seems to play an important role (Gomes et al., 2003) in bone growth onto these defects.

12.5

Cartilage tissue engineering using SPCL fiber-mesh scaffolds

Articular cartilage is responsible for the correct functioning of the skeleton, creating smooth gliding contact surfaces in the terminal parts of bones. Its main functions are related with shock absorbance, load bearing and reduction of surface friction in these parts. (Yaylaoglu et al., 1999). Articular cartilage’s

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ability to function as a weight bearing tissue is dependent on the two major components of its extracellular matrix, collagen type II and proteoglycans (Shepherd et al., 2004; Sikavitsas et al., 2001; Kon et al., 2000). Collagen is responsible for the tensile properties and the proteoglycans for compression resistance (Grodzinsky, 1983; Zhu et al., 2004). If this structural organization is affected by traumatic or degenerative conditions, pain and disability normally occur. Unlike other tissues such as bone for example, cartilage has a limited self repair capability. Moreover, even when some regeneration exists, fibrocartilage-like tissue is frequently formed that possesses lower mechanical properties, therefore generating a less functional tissue (Sikavitsas et al., 2001; Jeong Park et al., 2000; Gugala and Gogolewski, 2004). Due to this, different strategies have been suggested and studied to treat articular cartilage lesions and tissue engineering is one of those (Cancedda et al., 2003; Gugala and Gogolewski, 2004; Yaylaoglu et al., 1999). Most tissue engineering approaches are based on seeding cells in a polymeric matrix, made of synthetic or natural materials, that acts as a support for cells to proliferate and synthetize extracellular matrix (Hutmacher et al., 2007). Based on the promising results obtained previously in bone tissue engineering approaches, (Gomes et al., 2002; Tuzlakoglu et al., 2005) starch-polycaprolactone (SPCL) scaffolds were also put forward as an alternative in cartilage regeneration (Oliveira et al., 2006). Bovine articular chondrocytes were cultured on SPCL fibre scaffolds and characterized at different time periods using scanning electron microscopy (SEM), histological and immunological analysis. The scaffolds allowed an efficient penetration and colonization of the chondrocytes, while enabling extracellular matrix components to be deposited. Scanning electron micrographs (Fig. 12.1) showed that the bovine articular chondrocytes extensively colonized the scaffold structure, being widely present at the surface and penetrating the various pores. The morphology of the chondrocytes is one of normal and healthy cells (Barry and Murphy, 2004; Ciolfi et al., 2003), and these were forming multilayers. These observations indicate that with such an arrangement, the cells create a 3D environment which favours extracellular matrix formation. The constructs collected at four weeks are shown in Figures 12.1d-f. Higher cell coverage can be observed when compared with two weeks (a-c), indicating that the chondrocytes have proliferated during these periods. Figures 12.1g-i were taken from samples collected at six weeks and evidence extracellular matrix components deposition in the pericellular regions. Such statements will be confirmed later by histological and immunological analysis. Figure 12.2 shows optical microscopy images of different histological sections of scaffolds taken after four weeks (a) and six weeks (b-c) of culture. As previously observed in the SEM analysis, an increase in cell mass from four to six weeks can be observed. Moreover, a consistent adhesion interface with the SPCL fibres was frequently found during the analysis, as evidenced

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12.1 Scanning electron microscopy images of SPCL scaffolds seeded with bovine articular chondrocytes and cultured for 2 weeks (a–c), 4 weeks (d–f), and 6 weeks (g–i). © 2008, Woodhead Publishing Limited

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12.2 Optical microscopy images of histology sections obtained from SPCL scaffolds seeded with bovine articular chondrocytes and stained with hematoxylin-eosin. The images shown correspond to samples collected after 4 weeks (a), and 6 weeks (b–c) of culture.

in Figure 12.2.c (arrows). This fact suggests these tissue engineered constructs maintain their integrity in vivo, behaving as a single functional unit. Histology sections of SPCL scaffolds seeded with bovine articular chondrocytes were also stained with toluidine blue, a metachromatic stain that identifies glycosaminoglycans present in the extracellular matrix of hyaline cartilage (Figures 12.3a-c). A positive light purple staining was observed at both four and six weeks of culture, indicating that the chondrocytes had produced an extracellular matrix containing proteoglycans which is a good indicator towards the formation of a cartilage-like tissue. Collagen type II is the major protein produced by chondrocytes in articular cartilage, being involved in its weight bearing and adsorbing functions (Brodkin et al., 2004). Immunolocalization of collagen type II, and collagen type I for comparative analysis, was performed in sections obtained from SPCL seeded scaffolds after six weeks of culture (Figure 12.4). A difference in the expression pattern can be noted when comparing collagen type I and type II, with type II collagen displaying stronger antibody staining. Collagen type II is a good marker of tissue engineered hyaline-like cartilage (Kafienah and Sims, 2004) and therefore the detection of these proteins in the SPCL tissue engineered constructs is another indication of their hyaline-like nature.

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12.3 Optical microscopy images of histology sections obtained from SPCL scaffolds seeded with bovine articular chondrocytes stained with toluidine blue. The images shown correspond to samples collected after 4 weeks (a), and 6 weeks (b–c) of culture.

In conclusion the above mentioned studies demonstrate that SPCL scaffolds can support bovine articular chondrocytes adhesion, proliferation and differentiation, for up to six weeks of culturing. Chondrocytes were homogeneously distributed throughout the scaffolds’ structure while presenting good interface with the SPCL fibrous mesh. Toluidine blue staining and immunolocalization of collagens type I and type II evidenced glycosaminoglycans and collagen type II production, the two most important cartilage extracellular matrix components. Such results demonstrate that SPCL fibre based scaffolds constitute a suitable alternative for the regeneration of cartilaginous tissues, in addition to their promising performance in the bone tissue engineering area.

12.6

Advanced approaches using SPCL scaffolds for bone tissue engineering aiming at improved vascularization

Tissue engineering may be the ultimate solution to reconstruct and regenerate bone defects resulting from a number of clinical scenarios such as trauma, pathological degeneration, or congenital deformation (Cancedda et al., 2003; © 2008, Woodhead Publishing Limited

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12.4 Optical microscopy obtained from the immunolocalisation of collagen type I and type II in histology sections of SPCL scaffolds seeded with bovine articular chondrocytes. Images present results at 6 weeks for collagen type I, collagen type II, and normal goat serumcontrol (NGS).

Borenstein et al., 2002). Despite the major advances achieved within recent years, a critical obstacle in tissue engineering approaches, based on the in vitro culture of cell-scaffold constructs prior to implantation, is the ability to maintain large masses of living cells upon transfer from the in vitro culture conditions into the host (Wiesmann et al., 2004; Freed and Vunjak-Novakovic, 1998). All cells require access to substrate molecules (oxygen, glucose and amino acids) and a balance between consumption and local delivery of these substrates is needed if cells are to survive (Muschler et al., 2004). Therefore it is fundamental to have a functional vascular network supplying the implanted cell-construct and assuring the transport of oxygen and nutrients and clearance of catabolism products (Sieminski and Gooch, 2000; Nomi et al., 2002). In bone, vascularization is not only crucial to assure the metabolic survival of the implanted cells but it is also a critical process for osteogenesis and bone remodeling (Probst and Spiegel, 1997; Gerber and Ferrara, 2000; Choi et al., 2002). Therefore, the successful application of bone tissue engineering therapies in clinical practice is dependent on the development of new strategies that augment vascularization (Orban et al., 2002; Kannan et al., 2005). Some strategies have been suggested as an attempt to augment vascularization; nevertheless these are still far away from achieving the desire goal and © 2008, Woodhead Publishing Limited

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mainly focus on surgical techniques (Warnke et al., 2004; Warnke et al., 2006), growth factor delivery (Koch et al., 2006; Richardson et al., 2001) or cell-based therapies (Rouwkema et al., 2006; Choong et al., 2006), rather than 3D scaffolds with tailored surface and architecture.

12.6.1 Endothelialization of SPCL fiber-mesh scaffolds As it has been previously shown in this chapter, fiber-mesh scaffolds produced from a blend of starch with polycaprolactone (SPCL) are an excellent substrate for bone forming cells (Gomes et al., 2003; Gomes et al., 2006a; Gomes et al., 2006b). Even so, a successful implant for bone regeneration must also elicit the formation of a neovasculature supplying the scaffolding material (Soker et al., 2000). In this phenomenon of neovascularization, endothelial cells (ECs), the cell type that line the inner surface of blood vessels, play a critical role (Sumpio et al., 2002; Michiels, 2003). For this reason a scaffold for bone regeneration should have a proper surface and an architecture that promote an appropriate response from both seeded and host ECs. Previous in vitro studies have shown that ECs from both macro- and microvascular origin adhered to SPCL fiber-meshes and maintained their viability up to at least seven days, as assessed by calcein-AM (Santos et al., 2007). In addition, SPCL fiber-mesh scaffolds supported the maintenance of EC typical flattened morphology and other important functions such as endothelial integrity, which was maintained as shown by the expression of the intercellular junction proteins, platelet endothelial cell adhesion molecule-1 (PECAM-1) and vascular endothelial cadherin (VE-cadherin). The expression of these two cell-cell adhesion molecules is not only important for the stability of the monolayer, but also for the morphogenesis and physiology of the vessel wall (Cines et al., 1998). The participation of ECs in the inflammatory response through the expression of cell adhesion molecules is another function of this cell type (Muller, 2002). Indeed, when ECs growing on SPCL fiber-mesh scaffolds were exposed to a pro-inflammatory stimulus (lipopolysaccharide, LPS), an enhancement in the expression of cell adhesion molecules (CAMs), such as E-selectin and intercellular- and vascular cell adhesion molecules, was observed. This is a clear and a positive indication of the ability of ECs to participate in an inflammatory response. These overall results indicate that the surface chemical structure and topography of SPCL fiber-mesh scaffold are favorable for the maintenance of ECs viability, phenotype, morphology and functions.

12.6.2 Nano/micro fiber combined scaffold – innovative architecture Despite the positive and promising results that the study from Santos et al (2007) revealed, in science there is always room for improvement and

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Tuzlakoglu et al (2005) proposed a nano/micro fiber combined scaffold, a structure with an innovative architecture for bone regeneration based on SPCL fiber-mesh (Tuzlakoglu et al., 2005; Santos et al., 2007). The concept behind this novel scaffold is to mimic the biophysical structure of natural extracellular matrix (ECM) and to simultaneously combine two elements in the same construct: (a) microfiber-mesh aimed to give the mechanical support required during repair, obtained by a fiber-bonding process; and (b) nanonetwork that mimics ECM and aims to increase cell adhesion and motility, produced by electrospinning. This nano-network will provide an appropriate microenvironment for cells and have influence on cell functionality. Specifically regarding ECS, it was hypothesized that the existence of this structure in the nano-range could favor cellular adhesion between microfibers, and thus accelerating vascularization of the implanted scaffold. More details regarding the production of the scaffold can be found in the paper mentioned (Tuzlakoglu et al., 2005). Cell culture studies with human osteoblast-like cells and rat bone marrow cells demonstrated that, due to the presence of the nano-network, osteoblasts were able to bridge between microfibers and stretch themselves along the nanofibers, which resulted in a different cytoskeletal organization of the cells. In addition, alkaline phosphatase activity of the cells cultured on nano/micro fiber combined scaffold was higher than on control scaffolds, consisting of fiber-meshes without nanofibers (Tuzlakoglu et al., 2005). After proving the suitability of nano/micro fiber combined scaffolds for osteoblast differentiation and activity, another method of dealing with the interactions of developed scaffolds with ECs was pursued (Santos et al., 2008). The endothelial cells used in this work were of microvascular origin and this choice relied on its higher capacity to respond to angiogenic studies, to sprout and organize in tube-like structures (Dziubla and Lowman, 2004). The question marks raised in this work were related to the influence that this ECM-like architecture has on ECs growth, pattern, homo- (e.g. PECAM) and heterotypic interactions (e.g. CAMs) and on angiogenic potential. The nano/micro fiber combined scaffold was endothelialized after seven days of culture and ECs spanned between neighboring microfibers using the nanobridges formed by nano-fibers. Regarding the contact between adjacent cells through PECAM, immunofluorescence data revealed that even in the nano-range, cellular contact is maintained. In order to assess the angiogenic potential of nano/micro fiber combined scaffold, a special assay was performed, previously described, that reproduce in vitro the in vivo environment (Kirkpatrick et al., 2003). For that the scaffold was embedded into a collagen gel type I, whose function was to mimic the physical and functional properties of basement membrane, and the medium was supplemented with angiogenic growth factors (vascular endothelial growth factor, VEGF and basic fibroblast growth factor, bFGF). EC migration and spatial reorganization into tube-like structures was monitored by confocal microscopy of viable cells. In nano/

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micro fiber combined scaffolds, ECs migrated from the scaffolds into the gel and organized into tube-like structures. On the other hand, ECs growing on the control scaffold migrated mainly as a monolayer and fewer tube-like structures were detected. Perhaps the increased surface area as well as ECMlike structure of the nanonetwork facilitated EC migration and reorganization. Therefore nano/micro fiber combined scaffolds may enable the development of osteogenic scaffolds that promote the enhancement of neovascularisation of tissue engineered constructs.

12.7

Conclusions

The results described in this chapter show that SPCL scaffolds exhibit adequate porosity and mechanical properties to support cell adhesion and proliferation and also the tissue ingrowth upon implantation of the construct. An analysis of the degradation results showed that these starch based scaffolds are susceptible to enzymatic degradation, as detected by increased weight loss (within two weeks, the SPCL samples weight loss reached 20%) and decreased pH of the degradation solutions. With an increasing degradation time, the diameter of the SPCL fibers decreases significantly, increasing the porosity and consequently the available space for cells and tissue ingrowth during implantation time. In general, the described starch-based scaffolds allowed for the adhesion, proliferation and differentiation of marrow stromal cells towards the osteoblastic phenotype, under static and flow perfusion conditions. It was demonstrated that scaffold architecture, and especially pore interconnectivity, affect the homogeneity of the formed tissue. The work developed also emphasized the importance of the culturing system in bone tissue engineering approaches such as the one proposed in this study. Flow perfusion augments the functionality of scaffold/cell constructs grown in vitro as it combines both biological and mechanical factors that enhance cell differentiation and cell organization within the construct, towards the development of bone-like mineralized tissue. Additionally, this study also shows that flow perfusion bioreactor culture of marrow stromal cells combined with the use of appropriate starch based biodegradable scaffolds may constitute a promising approach for obtaining bone tissue substitutes and also provide a useful model to study bone formation and assess both in vitro and in vivo bone tissue engineering strategies. The in vivo studies showed that when SPCL scaffolds, alone or cell seeded, were orthotopically implanted in medium size animals such as goats, no significant inflammatory response was observed during the regeneration process, indicating their suitability for the in vivo environment. The above mentioned studies also demonstrate that SPCL scaffolds can support bovine articular chondrocytes adhesion, proliferation and differentiation, for up to six weeks of culturing. Chondrocytes were

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homogeneously distributed throughout the scaffold’s structure while presenting good interface with the SPCL fibrous mesh. Such results demonstrate that SPCL fiber based scaffolds constitute a suitable alternative for the regeneration of cartilaginous tissues, in addition to their promising performance in the bone tissue engineering area. Finally, results obtained so far demonstrate that it might be possible that the increased surface area as well as ECM-like structure of the nanonetwork in SPCL fiber meshes combining micro with nanofibers, facilitated EC migration and reorganization. The studies using nano/micro combined scaffolds also show that these belong to a new generation of scaffolds for bone regeneration whose 3D structure is tailored not only for osteogenesis but is also thought to meet strategies for the enhanced neovascularisation of such constructs.

12.8

Acknowledgments

The authors acknowledge the Portuguese Foundation for Science and Technology (FCT) for partial financial support through funds from the POCTI and/or FEDER programs and to the European Union funded STREP Project HIPPOCRATES (NMP3-CT-2003-505758). Most of the work described herein was carried out under the scope of the European NoE EXPERTISSUES (NMP3-CT-2004-500283).

12.9

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Thompson R, Yaszemski M and Mikos A (1997), Polymer Scaffold Processing. In Lanza R, Langer R and Chick W (Eds) Principles of Tissue Engineering. 1st edn. New York, Academic Press. Tietz N W and Shuey D F (1993), Lipase in Serum – the Elusive Enzyme – an Overview, Clinical Chemistry, 39, 746–756. Trojani C, Boukhechba F, Scimeca J C, Vandenbos F, Michiels J F, Daculsi G, Boileau P, Weiss P, Carle G F and Rochet N (2006), Ectopic bone formation using an injectable biphasic calcium phosphate/Si-HPMC hydrogel composite loaded with undifferentiated bone marrow stromal cells, Biomaterials, 27, 3256–3264. Tsuji H (2005), Poly(lactide) stereocomplexes: Formation, structure, properties, degradation, and applications, Macromolecular Bioscience, 5, 569–597. Tsuji H, Kidokoro Y and Mochizuki M (2006), Enzymatic degradation of biodegradable polyester composites of Poly(L-lactic acid) and poly(epsilon-caprolactone), Macromolecular Materials and Engineering, 291, 1245–1254. Tuzlakoglu K, Bolgen N, Salgado A J, Gomes M E, Piskin E and Reis R L (2005), Nanoand micro-fiber combined scaffolds: a new architecture for bone tissue engineering, J Mater Sci Mater Med, 16, 1099–1104. Vacanti C A and Bonassar L J (1999), An overview of tissue engineered bone, Clin Orthop, S375–381. Wake M, Patrick C and Mikos A G, Pore morphology effects on the fibrovascular tissue growth in porous polymers substrates, Cell Transplant, 3, 339, 1994. Wang Y, Uemura T, Dong J, Kojima H, Tanaka J and Tateishi T, Application of perfusion culture system improves in vitro and in vivo osteogenesis of bone marrow-derived osteoblastic cells in porous ceramic materials, Tissue Eng, 9, 1205, 2003. Warnke P H, Springer I N, Wiltfang J, Acil Y, Eufinger H, Wehmoller M, Russo P A, Bolte H, Sherry E, Behrens E and Terheyden H (2004), Growth and transplantation of a custom vascularised bone graft in a man, Lancet, 364, 766–770. Warnke P H, Wiltfang J, Springer I, Acil Y, Bolte H, Kosmahl M, Russo P A J, Sherry E, Lutzen U, Wolfart S and Terheyden H (2006), Man as living bioreactor: Fate of an exogenously prepared customized tissue-engineered mandible, Biomaterials, 27, 3163– 3167. Wiesmann H P, Joos U and Meyer U (2004), Biological and biophysical principles in extracorporal bone tissue engineering: Part II, International Journal of Oral and Maxillofacial Surgery, 33, 523–530. Williams J M, Adewunmi A, Schek R M, Flanagan C L, Krebsbach P H, Feinberg S E, Hollister S J and Das S (2005), Bone tissue engineering using polycaprolactone scaffolds fabricated via selective laser sintering, Biomaterials, 26, 4817–4827. Xiao Y, Qian H, Young W G and Bartold P M (2003), Tissue engineering for bone regeneration using differentiated alveolar bone cells in collagen scaffolds, Tissue Eng, 9, 1167–1177. Yaylaoglu M B, Yildiz C, Korkusuz F and Hasirci V (1999), A novel osteochondral implant, Biomaterials, 20, 1513–1520. Zhu Y B, Gao C Y, He T and Shen J C (2004), Endothelium regeneration on luminal surface of polyurethane vascular scaffold modified with diamine and covalently grafted with gelatin, Biomaterials, 25, 423–430.

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13 Chitosan-based scaffolds in orthopedic applications K. T U Z L A K O G L U and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

13.1

Introduction: Chemical and physical structure of chitosan and its derivatives

Chitosan is a biopolysaccaride obtained by a de-N-deacetylation process of chitin which is the primary structural polymer in arthropod exoskeletons and composed of N-acetyl-glucosamine and N-glucosamine units linked by β(1→4) glycosidic bonds. The deacetylation process is usually performed in the presence of hot alkali and it can result in chitosan with different molecular weights and degrees of deacetylation. Hence, the term ‘chitosan’ refers to a family of de-N-acetylated chitins with different degrees of deacetylation. Chitosan has three types of reactive groups, a primary amine group as well as both primary and secondary hydroxyl groups at C-2, C-3 and C-6 positions, respectively.1 Although all bring some important features to the polymer, the most important structural property of chitosan comes from the presence of the primary amine at the C-2″ position of the glucosamine residues, which is found in only few biological polymers in such a high content. Moreover, these amine groups are very reactive and can allow grafting of different substituents to chitosan as well as crosslinking the chitosan backbone to improve elasticity. The chemical properties and the charged state of chitosan change depending on the pH.2 At low pH values, the amines are protonated and positively OH NH2 O O

O O HO

O n

NH2 OH

13.1 Chemical structure of chitosan.

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charged, and chitosan behaves as a water soluble cationic polyelectrolyte. Conversely, these amines become deprotonated at high pH values, which makes chitosan insoluble in water. This soluble-insoluble transition of chitosan occurs at pH values between 6 and 6.5 which is a particularly suitable range for biological applications. When the pH reaches above 6.5, electrostatic repulsions between the polymeric chains are reduced and inter-polymer associations such as liquid crystalline domains or network junctions are formed. This characteristic of chitosan allows the production of fibers, films and 3D constructs from it. Another important property of chitosan is its biodegradability by lysozyme which is highly present in human body tissues and secretions.3 Depending on the degree of deacetylation and on the N-acetlyglucosamine group distribution along the polymer chains, it can completely degrade into small glycomino chains in vivo. Regarding in vivo biocompatibility of its degradation product, other workers have reported that these degradation products can be quickly eliminated by the kidney following intraperitoneal administration to mice, thus overcoming accumulation in the body.

13.2

Production methods for scaffolds based on chitosan and its composites or blends

13.2.1 Freeze-drying Freeze-drying is the most common and simple method to produce chitosanbased scaffolds. The scaffolds, with different pore size and porosity, can be formed by the simple procedure of freezing a chitosan solution in a suitable mould and subsequently lyophilizing the frozen structure. The freezing process provides the nucleation of ice crystals from solution and further growth along the lines of thermal gradients. Exclusion of the chitosan acetate salt from the ice crystal phase and subsequent ice removal by lyophilization generates a porous material. Average pore size of the scaffolds can be controlled by varying the freezing rate and hence the ice crystal size. Furthermore, the pore orientation is related to the geometry of the moulds and can be also controlled by changing thermal gradients during freezing. Madihally et al.4 reported that the pore size of chitosan scaffolds produced by lyophilization can be controlled within the range 1-250 µm, by varying the freezing conditions. They also demonstrated the influence of mould shape on the final microstructure of the pores. Scaffolds with radially oriented pores or parallel with a polygonal cross-section could be produced by this method using a cylindrical glass tube or a shallow dish as a mould, respectively. The freezing temperature and chitosan concentration are other important parameters that can have an effect on the final structure of the scaffold. An increase in the solution concentration can result in a larger mean pore size diameter due to the high

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ice crystal nucleation and growth. The freezing temperature used before lyophilization also controls the shape and size of the ice crystals. It has been demonstrated that the mean pore size diameter increases by decreasing the freezing temperature. 4 It was also reported that more uniform and interconnective pore structures can be obtained when lower freezing temperatures are used.5 In most cases, a secondary process is needed to develop more stable structures. The scaffolds produced only by lyophilization could have soluble chitosan acetate, which causes rapid swelling and ultimately dissolution in a neutral aqueous medium. This secondary process can be crosslinking before freezing, rehydration of samples in either sodium hydroxide (NaOH)4,5 or in an ethanol series,4 or a combination of crosslinking with rehydration.6,7

13.2.2 Freeze-gelling Although freeze drying is a very easy and efficient method to prepare porous structures from chitosan, the remove of residual solvent from the scaffolds is a time and energy consuming process and appears as a main drawback for this process. Besides this, if the temperature is not controlled well during the freeze-drying, another problem occurs related to the surface skin of the scaffold. When the temperature of the process is not low enough, the matrix will not be rigid enough to resist the interfacial tension caused by evaporation of the solvent. This causes a collapse in the porous structure and occurrence of a dense skin layer on the scaffold. To overcome these problems, recently solvent exchange phase separation has been proposed as an alternative method to the freeze-drying method to produce chitosan scaffolds with a similar structure.8–10 The idea of this process is to obtain the gelation of chitosan by using an alkaline environment below its gelation point. This process is also called ‘freeze-gelation’ since the gelation occurs below the freezing point of the polymer solution. Ho et al.9 prepared chitosan scaffolds with 80% porosity and pore size of 60–150 µm. The process was based on the immersion of a frozen chitosan solution in a precooled NaOH/ethanol solution in order to allow for the gelation of the chitosan. In a later study, the influence of freezing temperature and the concentration of acid and ethanol have been investigated by the same researchers.8 It has been reported that a higher freezing temperature and concentration of acetic acid in the scaffold solution increased the tensile stress and strain of the scaffold at maximum load, while ethanol concentration had a slight influence on the tensile stress. Chitosan/starch scaffolds have also been produced using a similar process.11 In a typical procedure, chitosan solution with starch in a dilute acetic acid solution has been placed in a plastic mould and frozen at 15–18°C overnight. It was then immersed in a precipitation solution consisting of NaOH and Na2SO4 in water12 or ammonia in ethanol.11 After precipitation, the samples

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were washed repeatedly with distilled water to remove excess salts until no pH changes were detected. The resulting scaffold showed an open, tubular, oriented pore structure with a porosity of about 80%11 or 5% depending on the ratio between chitosan and starch. The porosity of the scaffolds in the second study appeared to be low, where the idea was to obtain a porous structure by degradation of scaffolds with amylase and/or lipase in situ.

13.2.3 Wet spinning The wet spinning method is mostly used to produce natural fibers, such as chitin and chitosan fibers, which cannot be formed by either melt or dry spinning methods. Due to the strong inter-chain forces derived from the hydroxyl, acetamido and amino groups, chitin and chitosan tend to degrade at temperatures below their melting temperature. Therefore, melt spinning is typically not possible for chitin and chitosan. Besides that, these two natural polymers can only be dissolved in polar solvents which have high boiling points. As a consequence, dry spinning is also not practical for producing chitin and chitosan fibers. Wet spinning is the oldest spinning method to produce fibers from natural and synthetic polymers. A typical wet spinning unit consists of a viscose-type spinneret, a dope, a coagulation bath containing a coagulation solution, and rollers for winding, drawing and drying of filaments.13 Wet spinning of chitosan fibers can be obtained by extruding the viscous chitosan solution in dilute acid into a coagulation bath. The viscosity of the solution is an important parameter for the processibility of chitosan during extrusion. It also has an influence on the coagulation rate which includes the regeneration of the free amine form of chitosan. The coagulation bath must exhibit a high pH, such as aq. NaOH,14 aq. KOH,15 aq. NaOH-40% methanol,16 and aq. NaOH-Na2SO4 (or AcONa) mixture.17 In order to obtain fibers with good mechanical properties, some physical and chemical treatments (which are called drying treatments) can be used at the end of the process. It has been found that the drying treatments have a strong effect on the fiber properties as well as spinning conditions.18,19 In order to produce 3D structures from chitosan fibers, Tuzlakoglu et al. proposed a new simple methodology based on wet spinning.20 Figure 13.2 presents the route that was used for this purpose and the resulting 3D fiber structures. Briefly, a viscous chitosan solution was extruded into a coagulation bath containing 1M NaOH, 1M Na2SO4 and water in a certain ratio. After coagulation was completed, fibers were washed and subjected to a methanol treatment to improve mechanical properties of the scaffolds as was mentioned above. After drying, a scaffold 1 cm in diameter and 1 cm in height was obtained. These scaffolds have shown a swelling ability up to 170% in weight and

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Washing Alkaline bath

Treatment with methanol Moulding and drying at 60°C

13.2 Production route for 3D chitosan fiber mesh scaffold by wet spinning.

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Chitosan solution

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30% in volume in physiological conditions. Later studies also reported that these scaffolds could be coated with a bone-like apatite layer by a simple Bioglass® spraying methodology.21 The formed apatite layer on the chitosan fibers using this methodology can be seen in Figure 13.3. Denkbas et al.22 also used a wet spinning method to produce 3D chitosan fiber mesh scaffolds. In their method, ethanol was used as a coagulant in the presence of glutaraldehyde for crosslinking. The equilibrium swelling of the

(a)

(b)

13.3 Bone-like apatite layer on chitosan fiber mesh scaffolds, obtained by biomimetic coating. (a) × 100 (b) × 5000.

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scaffolds was found to be about 56.8%, achieved in about 30–90 min. Furthermore, they also investigated the release of 5-fluorouracil which had been loaded into the scaffold during processing. It has been reported that there was an influence of crosslinker density on the drug release rate.

13.2.4 Electrospinning Electrospinning is a relatively simple and efficient method to produce polymeric fibers on a nano scale. It has been used in polymer processing technology for more than 70 years and recently had much attention from the biomedical field, particularly in tissue engineering due to the structural properties of fabricated fibrous structures having diameters in the range close to the collagen fibers found in the natural extracellular matrix of about 30–130 nm.23 The main principle of electrospinning is to create an electrical field between a collector and a capillary connected with a reservoir containing a polymer solution in order to obtain nanofibers from a polymer solution or melt.24 This strong electrostatic field causes a deformation of a pendant droplet of the polymer solution at the capillary tip. Polymer droplet forms into a conical shape which is known as a Taylor cone. Once the applied voltage passes a threshold value, electrostatic forces overcome the surface tension and a fine charged jet is ejected. This charged polymeric jet then undergoes a stretching process which is accompanied by rapid solvent evaporation and the polymer forms a nanofiber when it reaches to the collector.25 Electrospinning has also been used to prepare nanofibrous mats from chitosan for different tissue engineering applications. However, electrospinning of pure chitosan has proven to be quite difficult.26,27 It has been claimed that the reason for this might be the ionic nature of chitosan.28 Under the high electrical field, the repulsive force between ionic groups within the polymer backbone is expected to inhibit the formation of continuous fibers during the electrospinning process. Therefore, the polymer forms ultrafine particles instead of ultrafine fibers. Up to date, there are only few reports on electrospining of pure chitosan. Ohkawa et al.27 succeed in producing chitosan nanofibrous mats with a fiber diameter about 330 nm by dissolving chitosan in trifluoroacetic acid. In another study, a concentrated acid solution (90%) has been used as a solvent for chitosan.29 It has been demonstrated that a uniform chitosan nanofibrous mat of average fiber diameter of 130 nm could be obtained from 7% chitosan in a concentred acetic acid solution. However, the difficulties in the removal of the solvents used becomes as a drawback in both studies. There are also many other reports on electrospinning chitosan fibers which have been in blends with other spinnable polymers, such as poly(ethylene oxide) (PEO),30,31 silk fibroin,32 poly(vinyl alcohol) (PVA),33 and collagen.34

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13.2.5 Rapid prototyping Rapid prototyping is a common name for a group of techniques, such as fused deposition modeling (FDM), laminated object manufacturing (LOM), three-dimensional printing (3DP), multiphase jet solidification (MJS) and 3D plotting, that can generate a physical model directly from computeraided design data. It is an additive process in which each part is constructed in a layer-by-layer manner.35 This technology allows one to produce a complex 3D structure of scaffolds with controlled architecture which means desired pore size, porosity and pore distribution. Some recent studies have been carried out to create chitosan scaffolds with a controlled architecture for tissue engineering. For instance, Ang and co-workers have used a new robotic desktop rapid prototyping system to design chitosan-hydroxyapatite (HA) scaffolds.36 This system consists of a computer-guided desktop robot and a one-component pneumatic dispenser. A chitosan/HA solution is extruded through a nozzle into a dispensing medium that contains sodium hydroxide and ethanol. After layer-by-layer deposition of chitosan in a preprogrammed pattern, the scaffolds are neutralized in sodium hydroxide solution in order to stabilize the final structure. It has also been demonstrated that it is possible to produce chitosan scaffolds with fully interconnected channel architecture. However, this system has the same disadvantage regarding the relation between the polymer concentration and precipitation rate. When the solution concentration is too high, precipitation occurs before it contacts the base layer, resulting in poor adhesion and failure to hold the strands down as they are dispensed. Conversely, the dispensed strands cannot precipitate fast enough to hold their shape in the medium when precipitation is too slow. To overcome those problems, this system has been improved by adding a dual dispensing part in the set-up.37 The scaffolds produced by this new system had an overall porosity of around 90% with a pore diameter of 200–500 µm. Rapid prototyping technology can also be used indirectly to fabricate chitosan based scaffolds. For instance, Manjubala et al.38 designed a dissolvable mould created by a rapid prototyping technique. This mould was then used for developing 3D chitosan/HA composite scaffolds. More recently, Jiankang et al.39 reported that it was also possible to design chitosan/gelatine scaffolds with a predefined architecture, using the same indirect rapid prototyping approach.

13.2.6 Particle aggregation Particle aggregation is another novel method for producing chitosan scaffolds. It was first reported by Malafaya et al. that chitosan scaffolds with well distributed and interconnected pores ranging from 100 to 400 µm can be obtained successfully by this method.40 The bioadhesive character of chitosan

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provides a strong bonding between the particles that leads to a very stable interface between the particles. These strong bonds result in very good mechanical properties (compressive modulus of around 300 MPa). However, the mean porosity of the scaffolds has been found to be around 30%, which is the main drawback of this process. More recently, Abdel-Fattah et al.41 also used a similar route to developed 3D chitosan scaffolds from chitosan microspheres. In their study, a diluted acetic acid solution was mixed with chitosan microspheres before moulding them into a stainless steel mould, while this kind of addition was not used in the first study mentioned above. Another difference between the two studies is the post-treatment method that they use after drying the scaffolds in the mould. In the first study, the researchers used glutaraldehyde as a crosslinker in order to obtain a more stable structure while the lyophilization technique was used for that in the second study. They were able to obtain chitosan scaffolds with a porosity of 19.2% and pore diameter of 199.6 µm. Related to these low porosity and pore diameter values, the compressive modulus of the scaffold was higher (662 MPa) than that in first study. Besides all the methods discussed above, melt based routes have been proposed by our research group to produce chitosan/polyester based scaffolds.42 Since chitosan degrades before melting, blends with poly(butylene succinate) (PBS), poly(butylene terephatalate adipate) (PBTA) and poly (ε-caprolactone) have been prepared to overcome this problem and make it suitable for melt based processing. It was possible to obtain porous scaffolds with interconnective pore structure by melt based compression moulding followed by salt leaching.

13.3

Orthopedic applications

13.3.1 Bone There are three main components to be considered in tissue engineering for bone regeneration: a scaffold, cells and the growth factors. The scaffold acts as a key component in this triangle since it serves as a template for cell interactions and the formation of a bone-extracellular matrix to provide structural support for the newly formed tissue.43 Due to this important role, an ideal scaffold should meet many criteria such as biodegradability, biocompatibility, appropriate surface characteristics, proper pore size and structure and mechanical properties similar to those of the bone repair site. Furthermore, it must be osteoconductive in order to support new bone formation from adjacent living bone. Osteoinduction is another important issue for rapid bone regeneration and can be simulated by delivery of some substitutes such as bone morphonogenetic proteins (BMPs), insulin-like growth factors (IGFs) and transforming growth factors (TGFs). Ideally, the scaffold should act as a delivery vehicle for those factors. Importantly, the scaffold should be

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a good host for the cells (osteoblasts or mesenchymal cells) that are seeded on it before implantation. Combining scaffolds with cells and growth factor is the most promising tissue engineering approach which allows rapid and effective bone regeneration in vivo in comparison to biomaterial matrices alone. Many synthetic and natural polymeric scaffolds, including chitosan and its derivatives, have been proposed for bone regeneration. Chitosan shows an ability to enhance wound healing rates, osteoconduction and provide antimicrobial properties. Moreover, due to its N-acetylglucosamine repeating units, chitosan has some similarity with glucosaminoglycans present in the natural extracellular matrix. This provides easy binding to growth factors such as fibroblast growth factor (FGF)44 which is normally present in the osseous trabecular tissue and endowed with mitogenic activity towards various types of mesenchymal cells, including osteoblasts.45 As discussed earlier, chitosan is a versatile polymer to be processed into different porous structures without the use of toxic solvents, which makes it an interesting material to be used as a scaffold for bone regeneration. Recently, an increasing number of anchorage-dependent cells, including bone cells, are being cultured on 3D chitosan based scaffolds for cell-based regenerative therapies.20,21,46–48 For instance, we have cultured human osteoblast-like cells on the chitosan fiber mesh scaffolds produced by wet spinning.20 The results demonstrated that the cells were able to adhere and proliferate on these scaffolds after seven days of culture. In a later study, we also evaluated with the same scaffolds the influence of a bone-like apatite layer coating on cell behaviour.21 When compared to the control samples (unmodified chitosan fiber mesh scaffolds) the cell population was found to be higher in the Ca-P biomimetic coated scaffolds, which indicates that the levels of cell proliferation on this kind of scaffold could be enhanced. Furthermore, it was also observed that the cells seeded in the Ca-P coated scaffolds had more spread and a flat morphology, which reveals an improvement in the cell adhesion patterns, phenomena that are always important in processes such as osteoconduction. The composites of chitosan with calcium phosphates have also been studied by many researchers and showed promising clinical application. Wang et al.49 investigated bone repair in radii and tibias of rabbits with phosphorylated chitosan reinforced calcium phosphate cements. Histological and histomorphological studies showed that P-chitosan-containing cements are biocompatible, bioabsorbable and osteoinductive. There was a progressive substitution at the interface between the implants and host bone. Furthermore, there was no adverse effect in tissue around the bone defects. Zhang et al.50 produced porous chitosan /HA/β-TCP composite scaffolds in order to provide high mechanical strength and a large surface/volume ratio for load-bearing bone repairing. An in vitro test with human osteoblast-like MG63 cells showed that the cells were able to be attached and proliferate on the surface of the composite

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scaffolds and migrated onto the pore walls. However, there was no significant difference in alkaline phosphatase (ALP) activity between the cells cultured on scaffolds and tissue culture dishes. In a later study, chitosan-calcium phosphate composite scaffolds were prepared by a solid-liquid phase separation method and tested with human osteoblast-like MG63 cells.51 It was found that ALP activity and osteocalcin production was significantly higher on the cells cultured on these scaffolds than those on chitosan scaffolds alone. Chitosan/tricalcium phosphate (TCP) composite sponges were also used for delivery of growth factors such as platelet-derived growth factor (PDGF).52 It was demonstrated that chitosan/ TCP sponges could promote osseous healing of the rat calvarial defects. Moreover, the addition of PDGF to these sponges can promote bone regeneration. Natural coralline, which is made of calcium carbonate (CaCO3), was also utilized as a gas foaming agent and reinforced material as well as osteoinductive component for preparation of chitosan based composites for bone repair.53 It was claimed that the coralline/chitosan composites, especially those having a high coralline content, may enhance adhesion, proliferation and osteogenic differentiation of mesenchymal stem cells (MSCs) in comparison with pure chitosan. Besides the composite of chitosan with bioactive ceramics, there are many reports on the use of chitosan blends with the other synthetic or natural polymers for bone regeneration. For instance, chitosan/collagen porous scaffolds with a pore size of 80–100 µm prepared by freeze drying were tested with a mouse osteoblast cell line (MC3T3-E1) for 21 days.54 It was demonstrated that these sponges could promoted growth and differentiation of osteoblasts into a mature stage. Li and co-workers55 constructed a scaffold made of chitosan and alginate with a porosity of around 92% and a compressive modulus of 8.16 MPa. They showed that osteoblasts seeded on these scaffold appeared to be attached, proliferated well and promoted the deposition of minerals in a very short time even without the use of an osteogenic medium. In a further step, it was demonstrated that chitosan/alginate scaffolds promoted rapid vascularization and deposited connective tissue and calcified matrix within the entire structure. Similar results were found with chitosan/gelatin/ chondroitin sulfate scaffolds prepared by freeze drying.56 The mesenchymal stem cells derived from rat bone marrow seeded on these scaffolds expressed osteopontin and osteocalcin after 21 days of culture in the presence of osteogenic medium. RGD (Arg-Gly-Asp) is the most attractive peptide sequence that is immobilized in the bone tissue engineering scaffolds to promote cell adhesion. Following this approach, Ho et al.57 modified chitosan scaffolds with RGDS (Arg-Gly-Asp-Ser) in order to evaluate the effect of immobilization on an osteoblast response in vitro. Experiments with rat osteosarcoma cells have shown that RGDS immobilization could enhance the attachment of these

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cells onto the chitosan and make the chitosan scaffolds more compatible for bone regeneration.

13.3.2 Cartilage Articular cartilage has a limited capacity for spontaneous healing because of the avascular nature of the tissue. Besides the many other traditional techniques, tissue engineering concepts have been proposed for the repair of articular cartilage. As in bone repair, the tissue engineering approach for articular cartilage involves the isolation of articular chondrocytes or their precursor cells that may be expanded in vitro and then seeded into a biocompatible matrix for cultivation and subsequent implantation into the joint. The loading of cartilage specific growth factors, such as transforming growth factor β1 (TGF-β1), or other bioactive agents (RGD, BMP7) can be also involved in this process. In articular cartilage, proteoglycans appear as one of the major macromolecules in its structure. These molecules consist of a core protein and covalently attached glycosaminoglycan (GAG) chains.58 GAGs are long, unbranched heteropolysaccharides, consisting of repeating disaccharide units, with the general structure: (uronic acid-amino sugar)n. It has been well known that GAGs can stimulate the chondrogenesis once the cartilage is damaged.59 Regarding their role, the use of GAGs or GAG analogs as components of cartilage tissue scaffolds appears to be a good approach for enhancing chondrogenesis. Chitosan, which has a similar characteristic as GAGs is one of the best candidates for that. Lahiji et al.60 showed that chitosan could support the expression of extracellular matrix proteins, collagen type II and aggrecan, in human chondrocytes. In this study, it was also found that chitosan has the ability to maintain round morphology of chondrocytes which is known to be indicative of a normal phenotypic characteristic of nondedifferentiated chondrocytes. 3D porous structures produced from chitosan by freeze drying have also been tested with porcine chondrocytes.61 The results indicated that chondrocytes grown on chitosan scaffolds synthesized an extracellular matrix containing proteoglycans and type II collagen after 18 days of culturing in a rotating-wall bioreactor. To improve the biological response to the scaffold, chitosan scaffolds containing other components of the natural extracellular matrix has been prepared by many researchers. For instance, hybrid fibers consisting of chitosan and hyaluronic acid, which is another major component of GAGs, have been produced by the wet spinning method and tested with articular chondrocytes isolated from rabbits.62 It was reported that the addition of hyaluronic acid could enhance cell adhesion and proliferation, as well as the synthesis of aggrecan. In another study, Chen et al.63 developed composite chondroitin-6-sulfate/dermatan sulphate/chitosan scaffolds for cartilage repair. As a different combination of the extracellular matrix components, collagen/chitosan/chondroitin sulphate scaffolds have

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been prepared by freeze drying and then combined with TGF-β1 loaded chitosan microspheres to enhance the cartilage formation.64 In another study, RGD containing protein was covalently bound to the chitosan-alginate-hyaluranon complexes to increase the cellular adhesion on chitosan.65 When the scaffolds with chondrocytes were implanted into a rabbit knee cartilage defect, a complete repair was observed after six months of implantation. Besides using RGD sequences, it was found that when BMP-7 was loaded, N, N-dicarboxymethyl chitosan scaffold induced repair of femoral articular cartilage lesions in rabbits.66

13.4

Conclusions and future trends

Despite extensive research on bone and cartilage conducted over recent decades, there is still a need of a definitive answer for successful repair of these tissues. It is our hope that tissue engineering approaches will provide a solution for repair of damaged bone and cartilage. At present, chitosan seems to be one of the most promising polymers for this purpose due to its several distinctive biological properties including good biocompatibility, biodegradability, and wound healing effect. Moreover, it can easily be formed into 3D porous structures which act as scaffolds in tissue engineering applications. However, there are still some drawbacks to be overcome such as insufficient mechanical properties and long degradation period. Combination of chitosan-based scaffolds with different polymers and ceramics and the use of new cell sources and growth factors will hopefully improve the success of chitosan in tissue engineering of bone and cartilage and allow its use in orthopedic clinical applications.

13.5

Acknowledgements

K. Tuzlakoglu thanks the Portuguese Foundation for Science and Technology for providing her a PhD scholarship (SFRH/BD/8502/2002). This work was partially supported by FCT Foundation for Science and Technology, through funds from the POCTI and/or FEDER programs and by the EU funded STREP Project HIPPOCRATES (NMP3-CT-2003-505758). This work was carried out under the scope of the European NoE EXPERTISSUES (NMP3CT-2004-500283).

13.6

References

1 Shi C, Zhu Y, Ran X, Wang M, Su Y and Cheng T, Therapeutic potential of chitosan and its derivatives in regenerative medicine, J Surgical Research, 2006, 133, 185– 192. 2 Yi H, Wu L, Bentley W E, Ghodssi R, Rubloff G W, Culver J N and Payne G F, Biofabrication with chitosan, Biomacromolecules, 2005, 6, 2881–2894.

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3 Varum K M, Myhr M M, Hjerde R J N and Smidsrod O, In vitro degradation rates of partially N-acetylated chitosans in human serum, Carbohydrate Research, 1997, 299(1-2), 99–101. 4 Madihally S V and Matthew H W T, Porous chitosan scaffolds for tissue engineering, Biomaterials, 1999, 20, 1133–1142. 5 Manjubala I, Scheler S, Bossert J and Jandt K D, Mineralisation of chitosan scaffolds with nano-apatite formation by double diffusion technique, Acta Biomaterialia, 2006, 2(1), 75–84. 6 Shen F, Cui Y L, Yang L F, Yao K D, Dong X H, Jia W Y and Shi H D, A study on the fabrication of porous chitosan/gelatin network scaffold for tissue engineering, Polymer International, 2000, 49(12), 1596–1599. 7 Zheng J P, Wang C Z, Wang X X, Wang H Y, Zhuang H and Yao K D, Preparation of biomimetic three-dimensional gelation/montmorillonite-chitosan scaffold for tissue engineering, React Funct Polym, 2007, 68, 780–788. 8 Hsieh C T S, Ho M, Wang D, Liu C, Hsieh C, Tseng H and Hsieh H, Analysis of freeze-gelation and crosslinking processes for preparing porous chitosan scaffolds, Carbohydrate Polym, 2007, 67, 124–132. 9 Ho M, Kuo P, Hsieh H, Hsien T, Hou L, Lai J and Wang D, Preparation of porous scaffolds by using freeze-extraction and freeze-gelation methods, Biomaterials, 2004, 25, 129–138. 10 Hsieh W, Chan C and Lin S, Morphology and characterization of 3D micro-porous structured chitosan scaffolds for tissue engineering, Colloids and Surfaces B: Biointerfaces, 2007, 57, 250–255. 11 Nakamatsu J, Torres F G, Troncoso O P, Min-Lin Y and Boccaccini A R, Processing and characterization of porous structures from chitosan and starch for tissue engineering scaffolds, Biomacromolecules, 2006, 7, 3345–3355. 12 Martins A M, Santos M I, Azevedo H S, Malafaya P B and Reis R L, Natural origin ‘smart’ scaffolds with in situ pore forming capability for bone tissue engineering applications, Biomaterials, 2007, (submitted). 13 Hirano S, Wet-spinning and applications of functional fibers based on chitin and chitosan, Macromol Symp, 2001, 168, 21–30. 14 Struszczyk H, Wawro D and Nicktaszcwicz A, Advances in Chitin and Chitosan, Elsevier: London, 1992, 580. 15 Knaul J Z, Hudson S M and Creber K A M, Improved mechanical properties of chitosan fibers, Journal of Applied Polymer Science, 1999, 72(13), 1721–1732. 16 Urbanczyk G W, Applications of Chitin and Chitosan, Technomic: Lancaster, 1997, 281. 17 Hirano S N K, Zhang M, Kim S K, Chung B G, Yoshikawa M and Midorikawa T, Chitosan staple fibers and their chemical modifications with some aldehydes, Carbohydrate Polym, 1999, 38, 293–298. 18 Knaul J Z, Hudson S M and Creber K A M, Improved mechanical properties of chitosan fibers, J Appl Poly Sci, 1999, 72, 1721–1732. 19 Knaul J, Hooper M, Chanyi C and Creber K A M, Improvements in the drying process for wet-spun chitosan fibers, Journal of Applied Polymer Science, 1998, 69(7), 1435–1444. 20 Tuzlakoglu K, Alves C M, Mano J F and Reis R L, Production and characterization of chitosan fibers and 3–D fiber mesh scaffolds for tissue engineering applications, Macromolecular Bioscience, 2004, 4(8), 811–819. 21 Tuzlakoglu K and Reis R L, Formation of bone-like apatite layer on chitosan fiber mesh scaffolds by a biomimetic spraying process, Journal of Materials ScienceMaterials in Medicine, 2007, 18(7), 1279–1286. © 2008, Woodhead Publishing Limited

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22 Denkbas E B, Seyyal M and Piskin E, Implantable 5-fluorouracil loaded chitosan scaffolds prepared by wet spinning, Journal of Membrane Science, 2000, 172(1-2), 33–38. 23 Matthews J A, Wnek G E, Simpson D G and Bowlin G L, Electrospinning of collagen nanofibers, Biomacromolecules, 2002, 3(2), 232–238. 24 Bognitzki M, Czado W, Frese T, Schaper A, Hellwig M, Steinhart M, Greiner A and Wendorff J H, Nanostructured fibers via electrospinning, Advanced Materials, 2001, 13(1), 70–72. 25 Li D and Xia Y N, Electrospinning of nanofibers: Reinventing the wheel, Advanced Materials, 2004, 16(14), 1151–1170. 26 Duan B, Dong C H, Yuan X Y and Yao K D, Electrospinning of chitosan solutions in acetic acid with poly(ethylene oxide), Journal of Biomaterials Science-Polymer Edition, 2004, 15(6), 797–811. 27 Ohkawa K, Cha D I, Kim H, Nishida A and Yamamoto H, Electrospinning of chitosan, Macromolecular Rapid Communications, 2004, 25(18), 1600–1605. 28 Min B M, Lee S W, Lim J N, You Y, Lee T S, Kang P H and Park W H, Chitin and chitosan nanofibers: electrospinning of chitin and deacetylation of chitin nanofibers, Polymer, 2004, 45(21), 7137–7142. 29 Geng X Y, Kwon O H and Jang J H, Electrospinning of chitosan dissolved in concentrated acetic acid solution, Biomaterials, 2005, 26(27), 5427–5432. 30 Bhattarai N, Edmondson D, Veiseh O, Matsen F A and Zhang M Q, Electrospun chitosan-based nanofibers and their cellular compatibility, Biomaterials, 2005, 26(31), 6176–6184. 31 Subramanian A, Vu D, Larsen G F and Lin H Y, Preparation and evaluation of the electrospun chitosan/PEO fibers for potential applications in cartilage tissue engineering, Journal of Biomaterials Science-Polymer Edition, 2005, 16(7), 861–873. 32 Park W H, Jeong L, Yoo D I and Hudson S, Effect of chitosan on morphology and conformation of electrospun silk fibroin nanofibers, Polymer, 2004, 45(21), 7151– 7157. 33 Li L and Hsieh Y L, Chitosan bicomponent nanofibers and nanoporous fibers, Carbohydrate Research, 2006, 341(3), 374–381. 34 Chen Z G, Mo X M and Qing F L, Electrospinning of collagen-chitosan complex, Materials Letters, 2007, 61(16), 3490–3494. 35 Yeong W Y, Chua C K, Leong K F and Chandrasekaran M, Rapid prototyping in tissue engineering: challenges and potential, Trends in Biotechnology, 2004, 22(12), 643–652. 36 Ang T H, Sultana F S A, Hutmacher D W, Wong Y S, Fuh J Y H, Mo X M, Loh H T, Burdet E and Teoh S H, Fabrication of 3D chitosan-hydroxyapatite scaffolds using a robotic dispensing system, Materials Science & Engineering C-Biomimetic and Supramolecular Systems, 2002, 20(1-2), 35–42. 37 Geng L, Feng W, Hutmacher D W, Wong Y S, Loh H T and Fuh J Y H, Direct writing of chitosan scaffolds using a robotic system, Rapid Prototyping Journal, 2005, 11(2), 90–97. 38 Manjubala I, Woesz A, Pilz C, Rumpler M, Fratzl-Zelman N, Roschger P, Stampfl J and Fratzl P, Biomimetic mineral-organic composite scaffolds with controlled internal architecture, Journal of Materials Science-Materials in Medicine, 2005, 16(12), 1111–1119. 39 Jiankang H D L, Yaxiong L, Bo Y, Bingheng L and Qin L, Fabrication and characterization of chitosan/gelatin porous scaffolds with predefined internal microstructures, Polymer, 2007, 48, 4578–4588. © 2008, Woodhead Publishing Limited

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40 Malafaya P B, Pedro A J, Peterbauer A, Gabriel C, Redl H and Reis R L, Chitosan particles agglomerated scaffolds for cartilage and osteochondral tissue engineering approaches with adipose tissue derived stem cells, Journal of Materials ScienceMaterials in Medicine, 2005, 16(12), 1077–1085. 41 Abdel-Fattah W I, Jiang T, El-Bassyouni G E T and Laurencin C T, Synthesis, characterization of chitosans and fabrication of sintered chitosan microsphere matrices for bone tissue engineering, Acta Biomaterialia, 2007, 3(4), 503–514. 42 Correlo V M, Boesel L F, Pinto A R, Bhattacharya M, Reis R L and Neves N M, ‘Novel 3–D chitosan/polyester scaffolds for bone and cartilage tissue engineering’, ESB 2005/19th European Conference on Biomaterials, Sorrento, Italy, 2005; Sorrento, Italy, 2005, 323. 43 Karageorgiou V, Kaplan D, Porosity of 3D biomaterial scaffolds and osteogenesis, Biomaterials, 2005, 26, 5474–5491. 44 Berscht P C, Nies B, Liebendorfer A and Kreuter J, Incorporation of basic fibroblast growth-factor into methylpyrrolidinone chitosan fleeces and determination of the invitro release characteristics, Biomaterials, 1994, 15(8), 593–600. 45 Hauschks PV, Growth factor effect in bone. In Bone, Hall B K (Ed) CRC Press: London, 1990; Vol. 1, pp 103–169. 46 Madihally S V and Matthew H W T, Porous chitosan scaffolds for tissue engineering, Biomaterials, 1999, 20(12), 1133–1142. 47 Zhao F, Yin Y J, Lu, W W, Leong J C, Zhang W J, Zhang J Y, Zhang M F and Yao K D, Preparation and histological evaluation of biomimetic three-dimensional hydroxyapatite/chitosan-gelatin network composite scaffolds, Biomaterials, 2002, 23(15), 3227–3234. 48 Seol Y J, Lee J Y, Park Y J, Lee Y M, Young-Ku Rhyu I C, Lee S J, Han S B and Chung C P, Chitosan sponges as tissue engineering scaffolds for bone formation, Biotechnology Letters, 2004, 26(13), 1037–1041. 49 Wang X H, Ma, J B, Wang Y N and He B L, Bone repair in radii and tibias of rabbits with phosphorylated chitosan reinforced calcium phosphate cements, Biomaterials, 2002, 23(21), 4167–4176. 50 Zhang Y and Zhang M Q, Three-dimensional macroporous calcium phosphate bioceramics with nested chitosan sponges for load-bearing bone implants, Journal of Biomedical Materials Research, 2002, 61(1), 1–8. 51 Zhang Y, Ni M, Zhang M Q and Ratner B, Calcium phosphate-chitosan composite scaffolds for bone tissue engineering, Tissue Engineering, 2003, 9(2), 337–345. 52 Lee Y M, Park Y J, Lee S J, Ku Y, Han S B, Klokkevold P R and Chung C P, The bone regenerative effect of platelet-derived growth factor-BB delivered with a chitosan/ tricalcium phosphate sponge carrier, Journal of Periodontology, 2000, 71(3), 418– 424. 53 Gravel M, Gross T, Vago R and Tabrizian M, Responses of mesenchymal stem cell to chitosan-coralline composites microstructured using coralline as gas forming agent, Biomaterials, 2006, 27(9), 1899–1906. 54 Arpornmaeklong P, Suwatwirote N, Pripatnanot P and Oungbho K, Growth and differentiation of mouse osteoblasts on chitosan-collagen sponges, International Journal of Oral and Maxillofacial Surgery, 2007, 36(4), 328–337. 55 Li Z S, Ramay H R, Hauch K D, Xiao D M and Zhang M Q, Chitosan-alginate hybrid scaffolds for bone tissue engineering, Biomaterials, 2005, 26(18), 3919– 3928. 56 Machado C B, Ventura J M G, Lemos A F, Ferreira J M F, Leite M F and Goes A M,

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3D chitosan-gelatin-chondroitin porous scaffold improves osteogenic differentiation of mesenchymal stem cells, Biomed Mater, 2007, 2, 124–131. Ho M H, Wang D M, Hsieh H J, Liu H C, Hsien T Y, Lai J Y and Hou L T, Preparation and characterization of RGD-immobilized chitosan scaffolds, Biomaterials, 2005, 26(16), 3197–3206. Suh J K F and Matthew H W T, Application of chitosan-based polysaccharide biomaterials in cartilage tissue engineering: a review, Biomaterials, 2000, 21(24), 2589–2598. Kosher R A, Lash J W and Minor R R, Environmental enhancement of in-vitro chondrogenesis.4. Stimulation of somite chondrogenesis by exogenous chondromucoprotein, Developmental Biology, 1973, 35(2), 210–220. Lahiji A, Sohrabi A, Hungerford D S and Frondoza C G, Chitosan supports the expression of extracellular matrix proteins in human osteoblasts and chondrocytes, Journal of Biomedical Materials Research, 2000, 51(4), 586–595. Nettles D L, Elder S H and Gilbert J A, Potential use of chitosan as a cell scaffold material for cartilage tissue engineering, Tissue Engineering, 2002, 8(6), 1009– 1016. Yamane S, Iwasaki N, Majima T, Funakoshi T Masuko T, Harada K, Minami A, Monde K and Nishimura S, Feasibility of chitosan-based hyaluronic acid hybrid biomaterial for a novel scaffold in cartilage tissue engineering, Biomaterials, 2005, 26(6), 611–619. Chen Y L, Lee H P, Chan H Y, Sung L Y, Chen H C and Hu Y C, Composite chondroitin-6-sulfate/dermatan sulfate/chitosan scaffolds for cartilage tissue engineering, Biomaterials, 2007, 28(14), 2294–2305. Lee J E, Kim K E, Kwon I C, Ahn H J, Lee S H, Cho H C, Kim H J, Seong S C and Lee M C, Effects of the controlled-released TGF-beta 1 from chitosan microspheres on chondrocytes cultured in a collagen/chitosan/glycosaminoglycan scaffold, Biomaterials, 2004, 25(18), 4163–4173. Hsu S H, Whu S W, Hsieh S C, Tsai C L, Chen D C and Tan T S, Evaluation of chitosan-alginate-hyaluronate complexes modified by an RGD-containing protein as tissue-engineering scaffolds for cartilage regeneration, Artificial Organs, 2004, 28(8), 693–703. Mattioli-Belmonte M, Gigante A, Muzzarelli R A A, Politano R, De Benedittis A, Specchia N, Buffa A, Biagini G and Greco F, N,N-dicarboxymethyl chitosan as delivery agent for bone morphogenetic protein in the repair of articular cartilage, Medical & Biological Engineering & Computing, 1999, 37(1), 130–134.

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14 Elastin-like systems for tissue engineering J. R O D R I G U E Z - C A B E L L O, A. R I B E I R O, J. R E G U E R A, A. G I R O T T I and A. T E S T E R A, Universidad de Valladolid, Spain

14.1

Introduction

Modern biomaterials science is characterized by a growing emphasis on identification of specific design parameters that are critical to performance, and by a growing appreciation of the need to integrate biomaterials design with new insights emerging from studies of cell-matrix interactions, cellular signalling processes, and developmental and systems biology.1 This leads us to a new concept, with an extraordinary use: ‘Nanobiotechnology’. The initial impetus for the development of technology at the nano-scale came from its relevance to electronics. However, interest has quickly been focused on biological systems. The search for common goals between this focus and biology arises from the idea that biology offers the most complex and sophisticated collection of functional nanostructures that exist. Understanding biological nanostructures will be enormously stimulating for nanotechnology; designers of nanomachines and biomolecular motors have much to learn from biological systems. Nanotechnology offers to biology new tools and biology offers nanotechnology access to new types of functional nanosystems (components of the cell) that are unquestionably interesting and possibly quite useful in the near future. Combined together, both could provide infinite possibilities for new designs in materials science.2 One of the most promising approaches in the fabrication of biomaterials is the bottom-up strategy, where materials are self-assembled, molecule by molecule, in order to perform new and hierarchically ordered complexes. This could be considered as a part of nanomaterials building, and a deep understanding of individual molecular behaviour and properties is required. For instance, the knowledge of the self-assembly of biological molecules provides an outstanding tool to obtain desired and functional supramolecular structures. Today, several self-assembled polymeric materials are well known and used as biomaterials. Natural biopolymers are an impressive example of advanced materials. Besides their unmatched biocompatibility and biodegradability, their bioactivity is one central feature of these biomaterials. 374 © 2008, Woodhead Publishing Limited

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Bioactive materials interact with biological systems and can interact with both the soft and hard tissues, promoting the healing and regeneration processes of surrounding tissues.

14.2

Genetic engineering of protein-based polymers

For many years, materials scientists have been investigating the possibilities of obtaining higher levels of control in polymer synthesis. Although important progress has been made, especially in recent years, the level of control of the characteristics of biomacromolecules was not reached using chemical synthesis. The development of molecular biology and genetic engineering techniques nowadays allow the design and bioproduction of polymers with absolute control of molecular architecture and the absence of randomness in the primary structure. Using these new molecular biology techniques, we are now able to obtain materials that mimic natural biomaterials, once we can create almost any DNA duplex coding in any amino-acid sequence. Many advantages can come from this approach. First, genetically-engineered protein-based polymers (GEPBPs)3 will, in principle, be able to show any simple or complex properties present in natural proteins. In this sense, this method offers an opportunity to exploit the huge resources, in terms of functionality, hoarded and refined to the extreme by biology during the long process of natural selection. GEPBPs can easily make use of the vast amount of functionality present in hundreds of thousands of different proteins existing in living organisms, from the smallest prions to viruses, for example. On the other hand, as we can construct coding gene, base by base, by following our own original designs and without being restricted to gene fragments found in living organisms, we can design and produce GEPBPs to obtain materials, systems and devices exhibiting functions of particular technological interest that are not displayed in living organisms. Moreover, from the point of view of a polymer chemist, the degree of control and complexity attained by genetic engineering is clearly superior to that achieved by even the most sophisticated polymer synthesis technologies. GEPBPs are characterized as being strictly monodisperse and can be obtained from a few hundred Daltons to more than 200 kDa; and this upper limit is continuously increasing.4 The number of different combinations attainable by combining the 20 natural amino acids is practically infinite. In a simple calculation, if we consider how many different combinations are possible to obtain small proteins consisting of, for example, 100 amino acids (their modest molecular mass would range in between 5.7 and 18.6 kDa), the figure is as high as 1.3 × 10130 possibilities, and the matter necessary to produce just a single copy of them would be 55 orders of magnitude higher that the whole known dark and luminous matter of the universe.3

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14.3

Genetic strategies for synthesis of proteinbased polymers

In synthesizing genes encoding repetitive protein based polymers, the techniques of molecular biology are typically employed to self-ligate monomer DNA fragments in a process of oligomerization. The monomer fragments must be oligomerized in a ‘head-to-tail’ orientation, and can be seamless in sequence or can contain intervening linkers between the desired repeats. Approaches to oligomerization can be broadly classified as either iterative, random or recursive, although these modes can be sequentially combined within the same implementation. Each of these methods is illustrated in Fig. 14.1. For iterative techniques, a DNA segment is oligomerized in a series of single, uniform steps; each step grows the oligomer by one unit length of the monomer gene. In random methods, an uncontrolled number of monomer DNA segments are oligomerized in a single step, creating a population of oligomerized clones of different lengths. This random approach of selfligation is referred to as ‘concatenation’ because the DNA segment is concatenated in a reaction that is analogous to the propagation step in chemical polymerization. Finally, in recursive approaches, DNA segments are joined E

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14.1 Schematics of three approaches to DNA oligomerization, (a) iterative, (b) random, (c) recursive.

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in sequential steps, with the length of the ligated segments growing geometrically in each step.5

14.4

State-of-the-art in genetically-engineered protein-based polymers (GEBPs)

Nowadays, genetic engineering of PBPs is still in its early infancy. The radically different approach in the methodology used to produce these polymers has resulted in the fact that, even now, a limited number of research groups and companies have made the effort to make this transition. Among these pioneer groups, the main interest has been mainly concentrated in two major polymer families, spider-silk like polymers6 and elastin-like polymers (ELPs), although some other interesting protein polymers have also been studied. Those include coiled-coil motifs and their related leucine zippers,7–9 β-sheetforming polymers, poly(allylglycine) and homopolypeptides such as poly(glutamic acid). Several families of polymers based on other elastic proteins such as resilin,10 abductin11 or gluten12 have attracted attention in recent years.2

14.5

Elastin-like polymers

ELPs are non-natural polypeptides composed of repeating sequences. They have their origin in the repeating sequences found in the mammalian elastic protein elastin that confers elasticity to structures such as skin and blood vessels. The most striking and longest sequence between cross-links in pig and cow is the undecapentapeptide (VPGVG)11.13,14 Along with this repeating sequence, others can be pointed out such as (VPGG) n , 15 (APGVGV)n.16 The importance of these polymers reside in the fact that they show a versatile and ample range of interesting properties that are difficult to find together in other materials, and that goes beyond their simple mechanical performance. Certainly, ELPs show a set of properties that places them in an excellent position towards designing advanced polymers for many different applications, including the most cutting edge biomedical and nanobiotechnological uses. Regarding their properties, some of their main characteristics are derived from the natural protein on which they are based. For example, the cross-linked matrices of these polymers retain most of the striking mechanical properties of elastin,17 i.e an almost ideal elasticity with Young’s modulus, elongation at break, etc. in the range of natural elastin and an outstanding resistance to fatigue.18,19 In addition, this mechanical performance is accompanied by an extraordinary biocompatibility.20 The polymer poly(VPGVG) is considered as a model for the ELPs. Most of the ELPs are based on the pentapeptide VPGVG (or its permutations),

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with amino-acid side-chains, excluding glycine, comprising simple aliphatic and mainly hydrophobic chains, without further functionalization. A wide variety of ELPs have been (bio)synthesized, based on the model ELP with the general formula VPGXG, where X represents any natural or modified amino acid except proline.21–23 All functional ELPs exhibit a reversible phase transition in response to changes in temperature,23 i.e. they show an acute thermo-responsive behaviour, associated with the existence of an inverse temperature transition (ITT). In aqueous solution, below the transition temperature Tt, the free polymer chains remain disordered, random coils in solution24 that are fully hydrated, mainly by hydrophobic hydration. This hydration is characterized by ordered clathrate-like water structures surrounding the apolar moieties of the polymer;25,26 structures somewhat similar to those described for crystalline gas hydrates,27 although more heterogeneous, and of varying perfection and stability. 26 However, above T t, the chain hydrophobically folds and assembles to form a phase-separated state of 63% water and 37% polymer in weight.28 In the folded state, the polymer chains adopt a dynamic, regular, non-random structure, called a β-spiral, involving type II β-turns as the main secondary feature, and stabilized by intra-spiral, inter-turn and inter-spiral hydrophobic contacts23 (see Fig. 14.2). This behaviour is the result of the ITT. In its folded and associated state, the chain loses essentially all of the ordered water structures of hydrophobic hydration.25 During the initial stages of polymer dehydration, hydrophobic association of the β-spirals takes on fibrillar form. This process starts from the formation of filaments composed of three-stranded dynamic polypeptide β-spirals that grow to several hundred nm before settling into a visible phase-separated state.23,29 The process of the ITT is completely reversible, and it goes back to the first state when the temperature is lowered below Tt.23 We have to keep in mind that, with this kind of polymer, and with the huge amounts of water playing an active role in increasing the dynamics of the

β-spiral formation

Association in twisted filaments

Aggregation

~1.8 nm ~5 nm

~1–2 µm

14.2 Mechanism for the inverse temperature transition of elastin-like polymers. From left to right: β-spiral formation of ELP molecules with three pentapeptides per turn, formation of twisted filaments or supercoil of β-spirals and their aggregation into microaggregates. Adapted with permission from J. Biomater. Sci. Polymer Edn.2

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polymer chain, regular polymer conformations cannot be understood as the fixed structures can be found, for example, in crystalline polymers. The folded state of the ELPs is characterized by a high chain mobility in which the β-spiral could be not the actual fixed structure but rather the ideal conformation around which the polymer chain oscillates.30 The transition can be easily followed either by turbidity measurements or by calorimetric methods, measuring the heat flow during the transition. The first method is characterized by a turbidity profile showing a sharp step, this increase in turbidity being caused by the formation of aggregates. Tt can be taken as the temperature at 10% or 50% change in the relative turbidity (Figure 14.3). In contrast, the differential scanning calorimetric (DSC) measurements are always characterized by a broad peak, expanding 20°C or more. In this case, Tt can be considered either as the onset of peak temperature; furthermore with this technique it is possible to obtain the enthalpy of the process as the area of the peak (Figure 14.3).30

14.6

Self-assembly behaviour of peptides and proteins

The key challenge in nanotechnology is to produce systems with desired functionalities at the nanometer scale. Over recent years, many synthetic strategies have been developed to obtain advanced nanodevices in an attempt to mimic the behaviour of biology in nature. Understanding the forces governing thermodynamic stability and specificity of the self-assembly events have become the central goal for biomimicking processes. Fabricating proteinbased devices through bottom-up approaches have been widely investigated since the preliminary studies of Drexler in 1981.31 Peptides and proteins are useful building blocks to obtain ordered nanostructures via self-assembly due to their well-stabilized folding, stability and protein–protein interactions.32 The attractive benefit of peptide self-assembly comes from the precise knowledge of their structural and functional information, allowing the construction of novel materials with tailored morphologies dictated by the individual building blocks.2

14.7

Self-assembly of elastin-like polymers (ELPs)

The self assembly of polymers is an emerging new field within material sciences, offering many potential applications in nanotechnological and nanobiotechnological fields. In relation to self-assembly, elastin undergoes a self-aggregation process in its natural environment. It is produced from a water soluble precursor, tropoelastin, which spontaneously aggregates into a covalently cross-linked fibrillar polymeric structure.33 The self-assembling ability of elastin resides in certain relatively short amino-acid sequences, as

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14.3 Thermal responsiveness of poly(VPGVG). (a) Turbidity profile as a function of temperature for a poly(VPGVG) 5 mg/L sample dissolved in water and DSC thermogram of a 20 µl (50 mg/L) aqueous solution of the same polymer (heating rate 5°C/min). (b) Photographs of aqueous solution (5 mg/mL) of this poly(VPGVG) below (5°C) and above (40°C) its Tt. Adapted with permission from Nanomedicine.30

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has been proposed by Yang et al.34 working on recombinant ELPs. The tendency to self-assemble in nanofiber conformation of ELPs can be extended to other topologies and nanostructural features.35–37 With the potential afforded by genetic engineering in designing new polymers, the growing understanding in the molecular behaviour of ELPs, and the enormous wealth of experimental and theoretical background gained during the past decade on the self-assembling characteristics of different types of block co-polymers, novel self-assembly properties are being unveiled within the ELP family. Reguera et al. showed that the ELP [(VPGVG)2-(VPGEG)-(VPGVG)2]15, previously found to be pH responsive, is able to form polymer sheets with self-assembled nanopores.37 AFM study of the topology of this ELP, containing equally spaced glutamicacid residues along the polymer chain, deposited by spin coating on a Si hydrophobic substrate at temperatures below Tt showed that in acid conditions, the deposited polymer presents a flat surface without any particular topological features. However, from neutral-basic solutions, the polymer monolayer clearly has an aperiodic pattern of nanopores (ca. 70 nm wide and separated by ca. 150 nm). This different behaviour as a function of pH has been explained in terms of the polarity of the free γ-carboxyl group of the glutamic acid. In the carboxylate form, this moiety shows a markedly higher polarity than the rest of the polymer domains and the substrate itself. Under this condition, the charged carboxylates impede any hydrophobic contact in their surroundings, which is the predominant mode of assembly for this kind of polymer. The charged domains, along with their hydration sphere, are then segregated from the hydrophobic surrounding giving rise to nanopore formation.2

14.8

Biocompatibility of ELPs

ELPs show an additional property which is highly relevant for the use of those polymers in advanced biomedical applications such as tissue engineering and controlled drug release. This is their tremendous biocompatibility. The complete series of the American Society for Testing Materials (ASTM) generic biological tests for materials and devices in contact with tissues, tissue fluids and blood demonstrate unmatched biocompatibility.20 Apparently the immune system just ignores these polymers because it cannot distinguish them from the natural elastin. Furthermore, monoclonal antibodies to the (GVGVP)n have not yet been produced, despite intensive efforts to do so.38 Incidentally, nowadays, it is believed that the high segmental mobility shown by the β-spiral, the common structural feature of ELPs, greatly helps in preventing the identification of these foreign proteins by the immune system.39 In addition, the secondary products of their bioabsorption are just simple and natural amino acids.

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14.9

Natural-based polymers for biomedical applications

Biomedical applications

Due to the new biomolecular techniques we are now capable of designing materials that will perform the same role in an organism as the natural ones. Here we can find the ELPs that, presenting a nice set of properties like smart-behaviour, self-assembly and extreme biocompatibility, are excellent for biomedical applications. This is applicable not only for ELPs that show a tailored bioactivity but also for the most basic ELPs; in this scenario we can find the polymers resulting from the repetition of the pentapeptides (VPGVG) and (VPGAG). Two different fields of interest have been targeted with ELPs in biomedical applications. The first application is for drug delivery systems, in which a polymer that can be in a matrix aggregate or device, releases drug gradually or when a certain stimulus triggers it. The second major application is the use of ELPs for tissue engineering in which the polymer, normally in the form of a matrix, works as a temporal scaffold that is gradually substituted by endogenous growing tissue. In many modern designs both functions, bioactive scaffolding and controlled release, come together to create a new generation of bioactive supports for cells.

14.10 ELPs for drug delivery ELP drug carriers can be divided into three different classes based on their intended end-use as carriers: (1) ELP homopolymers and pseudorandom copolymers for the systemic delivery of chemically conjugated radionuclides and chemotherapeutics;40 (2) block copolymers that thermally assemble into micelles or vesicles, designed for the encapsulation of hydrophobic drugs;41,42 and fusion proteins for the delivery of protein therapeutics.5 The first ELP-based drug delivery systems were reported by Urry. They were quite simple devices in which γ-radiated cross-linked poly(VPGVG) hydrogels of different shapes were loaded with a model water-soluble drug (Biebrich Scarlet).43 This drug was then released by diffusion. In this simple design, the extraordinary biocompatibility and the lack of pernicious compounds during the reabsorption of the device were exploited. The designs then became more complex. The basic VPGVG pentapeptide was functionalized by including some glutamic acids whose free carboxyl groups were used for cross-linking purposes. The cross-linker was of the type that forms carboxyamides, which were selected because of their ability to hydrolize at a given and controlled rate, releasing the polymer chains and, concurrently, any drug entrapped within the cross-linked slabs.44 Another example is the case of poly(VPAVG), were the self-association properties of the polymer were employed by our group in the development of various applications as controlled drug release devices. For example, Molina et al. 46 tested self-assembled nano- and micro-particles of

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poly(VPAVG), another version of ELP, as carriers of the model drug dexamethasone phosphate, in order to develop injectable systems for controlled drug release. In these particles, the drug is entrapped while the particles selfassemble as the temperature is raised above Tt. The slow diffusion of the drug has been considered as the main mechanism of drug delivery for this simple model, although more complex polymer designs based on this are possible. In a different approach, Betre et al. evaluated ELPs that aggregate below body temperature as a potential injectable depot for intra-articular drug delivery.46 Biodistribution studies revealed that the aggregating ELP has a 25-fold longer half-life in the injected joint than an equivalent molecular weight ELP that remains soluble and does not aggregate. These results suggest that the intra-articular delivery of ELP fusion proteins may be a viable strategy for the prolonged release of protein drugs for osteoarthritis.47 In an alternative approach, involving not only ELPs, injectable depots of silk-ELP hybrids were formed in situ for local delivery of DNA.48 This approach demonstrated the potential for sustained release of DNA from ELP depots, and may also be applicable for the release of other high molecular weight species such as proteins.47 The growing complexity of these systems is reaching nowadays levels that will lead us to almost perfect systems for controlled drug release. For example, Chilkoti’s group has produced nice examples of ELPs specially design for targeting and intracellular delivery, mainly in solid tumours. Taking advantage of the tunable Tt of ELPs they developed a non-invasive thermal targeting to solid tumours. This Tt when combined with hyperthermia treatment – the application of mild heat to the site of the tumour to promote the uptake of ELP – allows the drug to conjugate within the tumours.40,49,50 Actually, they showed, with in vivo studies of ELP delivery to two different types of implanted tumours in nude mice, that thermal targeting provided a nearly 23 fold increase in tumour localization when compared with non-heated controls.40,49 In summary they showed that, taking advantage of the Tt of the ELPs and local hyperthermia of tumours, it is possible to target the drugs, usually harmful for healthy cells, in a more precise and effective way.

14.11 Tissue engineering A tissue is composed of two major components: the cells and the extracellular matrix that the cells construct. Each tissue has a particular set of physical and chemical functions in fulfilling its role of sustaining the organ/or organism of which it is a part.51 The design of functional biomaterials that elicit cellular behaviour constitutes a major challenge for the fields of tissue engineering and materials science. Efforts to develop such materials have principally involved the design of scaffolds and hydrogels to mimic the dynamic interactions between cells and the extracellular matrix in vivo,52–54

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and the incorporation of extracelllar adhesion ligands and growth factors into engineered materials has proven effective in directing cellular responses in many applications.55–57 Knowing the central issues in tissue engineering, we can ask why sometimes materials that are used fail. Looking to all the processes that a mature or stem cell passes when it divides and spreads in growing tissue we clearly see that the cell is passing through the most vulnerable stage of its life cycle. And this is the main reason why materials that work in biomedical uses may fail when used for tissue engineering (the failure can be caused both by the material itself and by its biodegradation products). Additionally, we have to keep in mind that when we design a matrix for tissue engineering, we are trying to substitute for the natural extracellular matrix (ECM), at least transiently. Therefore, many aspects have to be taken into consideration in designing an adequate artificial ECM. Initially, the material developer must have a clear concern about the mechanical properties of the artificial scaffold. It is well know that, when properly attached to the ECM, cells sense the forces to which they are subjected via integrins. In this way, cells continuously sense their mechanical environment and respond by producing an ECM that adequately withstands the forces. In this sense, cells are very efficient force transducers. Therefore, any artificial ECM has to properly transmit forces from the environment to the growing tissue. Only in this way, can the new tissue build an adequate natural ECM that will eventually replace the artificial ECM. However, a stronger or too weak artificial ECM will cause its substitution by a too weak or too dense natural ECM, which can seriously compromise the success of tissue regeneration. In addition, growth tissues seem to need the input of mechanical deformation to create better structured tissues.38 The proteins of the natural ECM (fibronectin, collagen, elastin, etc.) contain in their sequences a huge number of bioactive peptides that are of crucial importance in the natural processes of wound healing. Those sequences include, of course, not only the well-known cell attachment sequences. In the natural ECM we find target domains for specific protease activity. Proteases, such as the metalloproteinases of the ECM, are only expressed and secreted to the extracellular medium when the tissue needs to remodel its ECM.58 They act on specific sequences that are present only in the proteins of the ECM, so they cannot cause damage to other proteins in their vicinity. It is also known that some fragments of these hydrolyzed ECM proteins, once released, show strong bioactivity, which includes the promotion of cell differentiation, spreading and angiogenesis, among other activities. Finally, growing tissue is delicately controlled by a well performed symphony of growth factors and other bioactive substances that are secreted by the cells. This is the scenario that tissue meets when passing through the difficult processes of growing and regenerating. One can say now, and after

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understanding how cells work when they are attaching and spreading in a scaffold, that the general aim of tissue engineering is to design temporary functional scaffolds that mimic essential elements of the structural and functional state of a tissue. Lots of materials have been applied in the field of tissue engineering, like the petroleum-based polymers and lately natural based polymers. The results are being improved with the passing of time but until today, and due to a lack of total understanding how cells work, the ideal material has not been found. This material has to be highly compatible; present excellent mechanical properties and include bioactive peptide sequences. A good candidate arises from this need, the ELPs. These polymers, besides presenting these requirements, are also smart materials that respond to changes in their environment. As a pioneer group in the developing of ELPs, Urry’s group was the first to develop systems based on these polymers for tissue engineering. Soon after they discovered the extraordinary biocompatibility of the (VPGVG)based ELPs,20 they tested the capability of these materials as raw materials for designing new scaffolds. The first candidates were simple VPGVG polymers and their cross-linked matrices. Surprisingly, tests on the cross-linked poly(VPGVG) matrices showed that cells did not adhere at all to this matrix and no fibrous capsules forms around it when it was implanted.59 Accordingly, this matrix and other states of the material have potential use in prevention of post-operative, post-trauma adhesions.59 Restoring an injured tissue to its normal state in many cases demands a means to prevent adhesion; that is, preventing abnormal bands of connective tissue binding tissues together inappropriately.51 In one of their works, Urry and his co-workers showed that when placed between an injured abdominal wall and an injured loop of bowel, a sheet of the cross-linked poly(GVGVP) prevented significant adhesion formation in 80% of the cases.60 Other applications of this basic polymer are in strabismus surgery, in preventing adhesion between rectus muscle and sclera of the eye;61 and in spinal surgery to correct herniated intervertebral discs.62 The absolute lack of cell adherence of the poly(GVGVP) could be seen as a drawback in the intended use of this kind of material for tissue engineering, but it is not. This polymer is ideal as a starting material since it provides adequate mechanical properties and biocompatibility and lacks unspecific bioactivity. Very soon these simple molecules were enriched with short peptides having specific bioactivities. Due to the polypeptide nature of the ELPs, those active short peptides were easily inserted within the polymer sequence. The first active peptides inserted in the polymer chain were the well-known general-purpose cell adhesion peptide RGD (R = L-arginine, G = glycine, D = L-aspartic acid)63 and REDV (E = L-glutamic acid, V = L-valine), which is specific for endothelial cells.64,65 Using ELPs with the RGD sequence, Urry et al. developed an excellent

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material for urological prostheses and intervertebral-disc restoration. In the first case they found that when using a matrix of the bioactive ELP and stimulating the matrix by filling and emptying, a greater density of bladder cells and extracellular matrix was obtained.38,66 In the second case, the selected bioelastic material was used to restore the proper dimensions and swelling pressure by injection into a disc previously depleted of its pulpous nucleus. The bioactive sequence was included to induce tissue regeneration while the disc is in an improved structural-functional state.51 Another interesting application of bioactive ELPs with cell adhesion sequences is the developing of vascular grafts. Severe atherosclerosis often requires surgical removal of the affected tissue and implantation of an autologous or synthetic vascular graft.67 The most widely used materials in synthetic vascular grafts, when used in small-diameter grafts, are characterized by high failure rates due to thrombosis and intimal hyperplasia.68–70 Tirrel’s group designed new vascular grafts attending two criteria, (a) enhancement of endothelial cell adhesion and (b) tuning the elastic modulus of the material to match that of the affected artery.71 For this, they developed two new ELPs, one containing the CS5 domain from the fibronectin and the other containing the well known RGD sequence. In both cases they substituted some of the valines by lysines for cross-linking purposes. Comparative studies of these two polymers showed that the grafts with the RGD cell-binding domained bound endothelial cells more strongly and elicited a faster cell spreading than the CS5 cell-binding domain.67 Using these ELPs they overwhelmed the two main problems in the vascular graft failure; the difference between veins’ tensile modulus and synthetic grafts and also the absence of a confluent endothelial monolayer.67 We can find in the literature other interesting work with ELPs as a raw material in tissue engineering applications. For example, one with a big potential regarding future applications is the work developed by Chilkoti’s group with genetically engineered ELPs for cartilaginous tissue repair. Cartilaginous tissues, such as articular cartilage, the meniscus, and intervertebral disc, contribute to mechanical load support, load distribution, and flexibility in joints of the body.72 These kinds of tissues virtually enable self-repair because they are avascular with a very low cell density. Taking advantage of the unique and favourable feature of the ELPs to form mechanical functional constructs in situ and their liquid-like behaviour below the temperature transition, Chilkoti’s group developed a method where they mixed chondrocyte cells with the ELPs in a way to create an injectable gel.72 In their studies they also demonstrated the ability of elastin-like polymers polypeptide gels to induce and support the chondrocytic differentiation of human adipose tissuederived adult stems cells in vitro in the absence of exogenous chondrogenic supplements, by promoting the expression of cartilage-specific genes and the accumulation of cartilage-specific extracellular matrix.46 This way

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they showed that ELPs have unique advantages as scaffolds for cartilage repair. There are currently examples based on more complex designs that include various bioactivities and other functionalities in an effort to mimic the complex composition and function of the natural extracellular matrix (ECM). Girotti et al. have bioproduced the ELP polymer depicted in Fig. 14.4.65 This ELP is made from a monomer 87 amino acids in length and has been produced with n = 10 (TE20109) (molecular mass 80.695 kDa). The monomer contains four different functional domains in order to achieve an adequate balance of mechanical and bioactive responses. First, the final matrix is designed to show a mechanical response comparable to the natural ECM, so that the matrix is produced over a base of an ELP of the type (VPGIG)n. This basic sequence assures the desired mechanical behaviour and outstanding biocompatibility. In addition, this basic composition endows the final polymer with smart and self-assembling capabilities, which are of high interest in the most advanced tissue engineering developments. The second building block is a variation of the first. It has a lysine substituting the isoleucine, so the lysine γ-amino group can be used for cross-linking purposes while retaining the properties of elastin-like polymers. The third domain is the CS5 human fibronectin domain. This contains the well-known endothelial cell attachment sequence, REDV, immersed in its natural sequence to potentiate its efficiency. Finally, the polymer also contains elastase target sequences to favour its bioprocessability by natural routes. The chosen elastase target sequence is the hexapeptide VGVAPG, which is found in natural elastin. The presence of this specific sequence in the artificial polymer guarantees that the polymer will be bioprocessed only when the growing tissue decides that it is time to substitute it by a natural ECM, while, in practice, it remains fully functional until that time. In addition, the activity of this domain is not restricted to being an inert target of protease activity. It is well known that these hexapeptides, as they are released by the protease action, have strong cell proliferation activity and other bioactivities related to tissue repair and healing.

n [(VPGIG)2–(VPGKG)–(VPGIG)2–(EEIQIGHIPREDVDYHLYP)–(VPGIG)2(VPGKG)–(VPGIG)2(VGVAPG)3]N

14.4 A schematic composition of the monomer of the ELP design described. The scheme shows the different functional domains of the monomer, which can be easily identified with their corresponding peptide sequences.

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14.12 Self-assembling properties of ELPs for tissue engineering One crucial issue in the interaction between cells and materials is the nanometric level. This interaction is essentially an interfacial interaction, in which the properties of the material surface play the key role. It is well known that certain topological features at the nanometric level can affect the performance of a given material; therefore, this phenomenon must be addressed and controlled. The surface self-assembly displayed by simple ELPs can also be present in more complex polymers such as those designed for tissue engineering. For example, this is the case for the ELP TE20109 shown previously. In this polymer, there are both predominantly apolar domains, such as those containing peptides (VPGIG) and (VGVAPG), and other domains, which can change their polarity by changing the pH, VPGKG and EEIQIGHIPREDVDYHLYP. Deposition of this polymer on an adequate substrate and at a suitable pH leads to a spontaneous nanostructuration of the polymer surface. This can be seen in Figure 14.5a-c.

14.13 Processability of ELPs for tissue engineering Of course, those protein-based materials can also be processed as conventional tissue engineering materials. Several examples can be found in the literature. The way in which matrices and hydrogels from these polymers can be prepared is diverse; for example, by γ-irradiation4 or by placing specific amino acids with functional groups for easy further cross-linking.4,19, 73–75 For instance, by using the free γ-amino group of lysine and hexamethylene diisocyanate as a cross-linker, the polymer His-Tag-{[(VPGIG)2-VPGKG-(VPGIG)2AVTGRGDSPASS-(VPGIG)2-VPGKG-(VPGIG)2]6} (TE20211), forms stable hydrogels (Figure 14.6).76

730 nm (a)

(b)

(c)

14.5 Atomic force microscope images of the deposition of a polymer for tissue engineering with different conditions. (a) Polymer solution at pH 7 deposited on hydrophobic silicon substrate. (b) Polymer solution at pH 12 deposited on hydrophobic silicone substrate. (c) Polymer solution at pH 12 deposited on hydrophilic mica.

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14.6 Cross-linked matrix of the RGD-containing polymer TE20211. Reproduced with permission from Nanomedicine.30

Additionally, in this cross-linked state, the polymer still retains its smart nature. In this sense, the matrix tends to hydrate by hydrophobic hydration below its Tt while it shrinks and de-swells above Tt . Interestingly, in the swollen state (below Tt), the weight of the hydrogel shown in Figure 14.6 is approximately 2 g, and only 20 mg of polymer have been used to form the matrix. Therefore, the ELPs can be considered as exceptional superabsorbents. Also, the ITT can still be observed by DSC. Of course, the microstructures and properties are different in both states. Below Tt, the matrix shows a honeycomb porous structure that is lost at temperatures above Tt (Figure 14.7). That greatly affects, among many other properties, its mechanical and diffusion properties. Thinking in terms of 3D tissue engineering, the porous structure found at temperatures below Tt is not wide enough to permit the cells to colonize the interior of the matrix. However, many well established techniques for the formation of pores can be used to achieve a more convenient porous scaffold. For example, by using dimethylformamide (DMF) as the solvent during cross-linking, a salt (NaCl) leaching method can be used to obtain adequate porous scaffolds (Figure 14.8).30

14.14 Future trends The conjugation of nanotechnological concepts and the know-how of biological processes leads us to a step where we are able to understand the bases of the

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10 µm

20 µm (a)

(b)

14.7 Scanning electron microscope micrographs of a TE20211 crosslinked hydrogel (a) above 37°C and (b) below 15°C, its Tt. Reproduced with permission from Nanomedicine.30

20 µm

14.8 Scanning electron micrograph of a TE20211 cross-linked porous hydrogel obtained by SALT leaching as explained in the text. Reproduced with permission from Nanomedicine.30

design of new materials. Although nanobiotechnology is in its childhood, the current work is laying the foundations for the future on materials science in tissue engineering and other areas. Since the potential of genetic-engineering techniques has not been fully exploited we can look at this tool as unlimited. The design of new materials is getting closer and closer to the performance

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of materials and molecular machines found in nature. Due to the close collaboration between biology, medicine and nanotechnology, we can now assume that in the near future we will be witnesses to the appearance of a new generation of biomaterials with astonishing properties, as they will progressively integrate the new concepts that come out every day in the most diverse areas. Taking into account all the recent applications described previously and their interesting and unconventional properties, the elastin-like polymers will play a key role in this new generation of materials.

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14 Yeh H, Ornstein-Goldstein N, Indik Z, Sheppard P, Anderson N, Rosenbloom J C, Cicilia G C, Yoon K and Rosenbloom J (1987), Sequence variation of bovine elastin mRNA due to alternative splicing, Collagen Rel Res, 7(4), 235–247. 15 Sandberg L B, Soskel N T and Leslie J G (1981), Elastin structure, biosynthesis, and relation to disease states, New Engl J Med, 304(10), 566–579. 16 Indik Z, Yeh H, Ornstein-Goldstein N, Sheppard P, Anderson N, Peltonen L, Rosenbloom J C and Rosenbloom J (1987), Alternative splicing of human elastin mRNA indicated by sequence analysis of cloned genomic and complementary DNA, Proc Natl Acad Sci, 84, 5680–5684. 17 Ayad S, Humphries M, Boot-Handford R, Kadler K and Shuttleworth A (1994), The Extracellular Matrix Facts Book, San Diego, CA, Academic Press. 18 Urry D W, Luan C H, Harris C M and Parker T (1997), ‘Protein-based materials with a profound range of properties and applications: the elastin Tt hydrophobic paradigm’, in McGrath K, Kaplan D, Protein Based Materials, Boston, MA, Birkhauser, 133– 177. 19 Di Zio K and Tirrell D A (2003), Mechanical properties of artificial protein matrices engineered for control of cell and tissue behavior, Macromolecules, 36, 1553–1558. 20 Urry D W, Parker T M, Reid M C and Gowda D C (1991), Biocompatibility of the Bioelastic Materials, Poly(GVGVP) and Its Gamma-Irradiation Cross-Linked Matrix – Summary of Generic Biological Test-Results, J Biact Compat Polym, 6, 263–282. 21 Gowda D C, Parker T M, Harris R D and Urry D W (1994), in Peptides: Design, Synthesis and Biological Activity, Basava C and Anantharamaiah G M (Eds), Boston M A, Birkhäuser, p 81. 22 Martino M, Perri T and Tamburro A M (2002), Biopolymers and biomaterials based on elastomeric proteins, Macromol Biosci, 2, 319–328. 23 Urry D W (1993), Molecular machines: how motion and other functions of living organisms can result from reversible chemical changes, Angew Chem Int Edn, 32, 819–841. 24 San Biagio P L, Madonia F, Trapane T L and Urry D W (1988), Overlap of elastomeric polypeptide coils in solution required for single phase initiation of elastogenesis, Chem Phys Lett, 145, 571–574. 25 Urry D W (1997), Physical chemistry of biological free energy transduction as demonstrated by elastic protein-based polymers, J Phys Chem B, 101, 11007–11028. 26 Rodríguez-Cabello J C, Alonso M, Perez T and Herguedas M M (2000), Differential scanning calorimetry study of the hydrophobic hydration of the elastin-based polypentapeptide, poly(VPGVG), from deficiency to excess of water, Biopolymers, 54(4), 282–288. 27 Pauling L and Marsh R E (1952), The structure of chlorine hydrate, Proc Natl Acad Sci USA, 38, 112–118. 28 Urry D W, Trapane T L and Prasad K U (1985), Phase-structure transitions of the elastin polypentapeptide-water system within the framework of compositiontemperature studies, Biopolymers, 24, 2345–2356. 29 Manno M, Emanuele A, Martorana V, San Biagio P L, Bulone D, Palma-Vittorelli M B, McPherson D T, Xu J, Parker T M and Urry D W (2001), Interaction of processes on different length scales in a bioelastomer capable of performing energy conversion, Biopolymers, 59, 51–64. 30 Rodriguez-Cabello J C, Prieto S, Arias F J, Reguera J and Ribeiro A (2006), Nanobiotechnological approach to engineered biomaterial design: the example of elastin-like polymers, Nanomedicine, 1(3), 267–280.

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31 Drexler K E (1981), Molecular engineering: An approach to the development of general capabilities for molecular manipulation, Proc Natl Acad Sci USA, 78, 5275– 5278. 32 Rajagopal K and Schneider J P (2004), Self-assembling peptides and proteins for nanotechnological applications, Curr Opin Struct Biol, 14, 480–486. 33 Miao M, Cirulis J T, Lee S and Keeley F W (2005), Structural Determinants of Cross-linking and Hydrophobic Domains for Self-Assembly of Elastin-like Polypeptides, Biochemistry, 44, 14367–14375. 34 Yang G C, Woodhouse K A and Yip C M (2002), Substrate-Facilitated Assembly of Elastin-Like Peptides: Studies by Variable-Temperature in Situ Atomic Force Microscopy, J Am Chem. Soc, 124, 10648–10649. 35 Wright E R and Conticello V P (2002), Self-assembly of block copolymers derived from elastin-mimetic polypeptide sequences, Adv Drug Deliv Rev, 54, 1057–1073. 36 Lee T A T, Cooper A, Apkarian R P and Conticello V P (2000), Thermo-reversible self-assembly of nanoparticles derived from elastin-mimetic Polypeptides, Adv Mater, 12, 1105–1110. 37 Reguera J, Fahmi A, Moriarty P, Girotti A and Rodríguez-Cabello J C (2004), Nanopore formation by self-assembly of the model genetically engineered elastin-like polymer [(VPGVG)2(VPGEG)(VPGVG)2]15, J Am Chem Soc, 126, 13212– 13213. 38 Urry D W et al. (1997), in Domb A, Kost J and Wiseman D, Handbook of Biodegradable Polymers, Chur, Harwood Academic Publishers, 367–386. 39 Urry D W (2005), What Sustains Life? Consilient Mechanism for Protein-based Machines and Materials, NY, Springer-Verlag. 40 Meyer D E, Kong G A, Dewhirst M W, Zalutsky M R and Chilkoti A (2001), Targeting a genetically engineered elastin-like polypeptide to solid tumors by local hyperthermia. Cancer Res, 61(4), 1548–1554. 41 Chung J E, Yokoyama M, Yamato M, Aoyagi T, Sakurai Y and Okano T (1999), Thermo-responsive drug delivery from polymeric micelles constructed using block copolymers of poly(n-isopropylacrylamide) and poly(butylmethacrylate), J Controlled Release, 62(1-2), 115–127. 42 Chung J E, Yokoyama M, Aoyagi T, Sakurai Y and Okano T (1998), Effect of molecular architecture of hydrophobically modified poly(n-isopropylacrylamide) on the formation of thermoresponsive core-shell micellar drug carriers, J Controlled Release, 53, 119–130. 43 Urry D W, Gowda D C, Harris C M and Harris R D (1994), ‘Bioelastic materials and the ∆Tt-Mechanism in drug delivery’, in: Ottenbrite R M, Polymeric Drugs and Drug Administration, Washington DC, CRC Press, 15. 44 Urry D W (1990), Preprogrammed drug delivery systems using chemical triggers for drug release by mechanochemical coupling, Polym Mater Sci Eng, 63, 329–336. 45 Herrero-Vanrell R, Rincón A C, Alonso M, Reboto V, Molina-Martinez I T and Rodriguez-Cabello J C (2005), Self-assembled particles of an elastin-like polymer as vehicles for controlled drug release, J Control Rel, 102(1), 113–122. 46 Betre H, Ong S R, Guilak F, Chilkoti A, Fermor B and Setton L A, (2006), Chondrocytic differentiation of human adipose-derived adult stem cells in elastin-like polypeptide, Biomaterials, 27, 91–99. 47 Chilkoti A, Christensen T and MacKay J A (2006), Stimulus responsive elastin biopolymers: applications in medicine and biotechnology, Current Opinion in Chemical Biology, 10, 1–6.

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48 Megeed Z, Haider M, Li D, O’Malley J B W, Cappello J and Ghandehari H (2001), In vitro and in vivo evaluation of recombinant silk-elastinlike hydrogels for cancer gene therapy, J Control Release, 94, 433–445. 49 Chilkoti A, Dreher M R, Meyer D E and Raucher D (2002), Targeted drug delivery by thermally responsive polymers, Advanced Drug Delivery Reviews, 54, 613–630. 50 Meyer D E, Shin B C, Kong G A, Dewhirst M W and Chilkoti A (2001), Drug targeting using thermally responsive polymers and local hyperthermia, Journal of Controlled Release, 74, 213–224. 51 Urry D W (1999), Elastic molecular machines in metabolism and soft tissue restoration, TIBTECH, 17, 249–257. 52 Kleinman H K, Philp D and Hoffman M (2003), Role of the extracellular matrix in morphogenesis, Curr Opin Biotechnol, 14, 526–532. 53 Hubbell J A (2003), Materials as morphogenetic guides in tissue engineering, Curr Opin Biotechnol, 14, 551–558. 54 Maskarinec S A and Tirrell D A (2005), Protein engineering approaches to biomaterials design, Curr Opin in Biotechnol, 16, 1–5. 55 Griffith L G and Naughton G (2002), Tissue engineering – current challenges and expanding opportunities, Science, 295, 1009–1014. 56 Hirano Y and Mooney D J (2004), Peptide and protein presenting materials for tissue engineering, Adv Mater, 16(1), 17–25. 57 Sakiyama-Elbert S E and Hubbell J A (2001), Functional biomaterials: design of novel biomaterials, Annu Rev Mater Res, 31, 183–201. 58 Sternlicht M D and Werb Z (2001), How matrix metalloproteinases regulate cell behavior, Annu Rev Cell Dev Biol, 17, 463–516. 59 Urry D W, Nicol A, Gowda D C, Hoban L D, McKee A and Williams T (1993), ‘Medical applications of bioelastic materials’ in: Gebelein C G, Biotechnological Polymers: Medical, Pharmaceutical and Industrial Applications, Atlanta, GA: Technomic Publishing Co. Inc., 82–103. 60 Hoban L D, Pierce M, Quance J, Hayward McKee A and Gowda D C (1994), The use of polypenta-peptides of elastin in the prevention of postoperative adhesions, J Surg Res, 56, 179–183 61 Elsas F J, Gowda D C and Urry D W (1992), Synthetic polypeptide sleeve for strabismus surgery, J Pediatr Ophthalmol Strabismus, 29, 284–286. 62 Alkalay R N, Kim D H, Urry D W, Xu J, Parker T M and Glazer P A (2003), Prevention of post-laminectomy epidural fibrosis using bioelastic materials, Spine, 28, 1659–1665. 63 Urry D W, Pattanaik A, Xu J, Woods T C, McPherson D T and Parker T M (1998), Elastic protein-based polymers in soft tissue augumentation and generation, Journal of Biomaterials Science-Polymer Edition, 9(10), 1015–1048. 64 Panitch A, Yamaoka T, Fournier M J, Mason T L, Tirrell D A (1999), Design and biosynthesis of elastin-like artificial extracellular matrix proteins containing periodically spaced fibronectin CS5 domains, Macromolecules, 32, 1701–1703. 65 Girotti A, Reguera J, Rodríguez-Cabello J C, Arias F J, Alonso M and Testera A M (2004), Design and bioproduction of a recombinant multi(bio)functional elastin-like polymer containing cell adhesion sequences for tissue engineering purposes, J Mater Sci- Mater Med, 15, 479–484. 66 Urry D W and Pattanaik A (1997), Elastic protein-based materials in tissue reconstruction, Ann New York Acad Sci, 831, 32–46. 67 Liu J C, Heilshorn S C and Tirrell D A (2004), Comparative cell response to artificial

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extracellular matrix proteins containing the RGD and CS5 cell-binding domains, Biomacromolecules, 5, 497–504. Nugent H M and Edelman E R (2003), Tissue engineering therapy for cardiovascular disease, Circ Res, 92, 1068–1078. Salacinski H J, Tiwari A, Hamilton G and Seifalian A M (2001), Cellular engineering of vascular bypass grafts: role of chemical coatings for enhancing endothelial cell attachment, Med Biol Eng Comput, 39, 609–618. Bos G W, Poot A A, Beugeling T, van Aken W G and Feijen J (1998), Small-diameter vascular graft prostheses: current status, Arch Physiol Biochem, 106, 100–115. Heilshorn S C, Liu J C and Tirrell D A (2005), Cell-binding domain context affects cell behavior on engineered proteins, Biomacromolecules, 6, 318–323. Betrea H, Setton L A, Meyer D E and Chilkoti A (2002), Characterization of a genetically engineered elastin-like polypeptide for cartilaginous tissue repair, Biomacromolecules, 3, 910–916. Welsh E R and Tirrel D A (2000), Engineering the extracellular matrix: a novel approach to polymeic biomaterials.I. Control of the physical properties of artificial protein matrices designed to support adhesion of vascular endothelial cells, Biomacromolecules, 1, 23–30. Trabbi-Carlson K, Setton L A and Chilkoti A (2003), Swelling and mechanical behaviors of chemically cross-linked hydrogels of elastin-like polypeptides, Biomacromolecules, 4, 572–580. McMillan R A, Caran K L, Apkarian R P and Conticello V P (1999) High-resolution topographic imaging of environmentally responsive elastin-mimetic hydrogels. Macromolecules, 32, 9067–9070. Prieto S, Espirito-Santo V, Alonso M, Testera A M, Mano J and Rodriguez-Cabello J C, Physical properties of artificial extracellular matrix of a cross-linked elastin-like polymer designed for tissue engineering purposes. (in preparation).

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15 Collagen-based scaffolds G. C H E N, N. K A WA Z O E and T. T AT E I S H I, National Institute for Materials Science, Japan

15.1

Introduction

Collagen is the most abundant and ubiquitous protein in vertebrates. More than 20 genetically distinct collagens have been identified. Collagen molecules are obtained by digesting the insoluble collagen tissue with pepsin to cleave the crosslinking sites of the tissue. A collagen solution is obtained by dissolving the collagen molecules in an aqueous solution. The versatile properties of collagen have made it one of the most useful biomaterials for tissue engineering. It can be in the form and shape of natural tissue, where other molecular and cellular components have been removed by biochemical treatments, while retaining the crosslinked structure of the tissue. Alternatively, it can be reconstituted in the form of porous scaffolds and gels from a collagen solution. Collagen has also been hybridized with other naturally derived polymers, biodegradable synthetic polymers, and inorganic biomaterials to improve their mechanical properties and biocompatibility. Collagen-based scaffolds have been widely used for tissue engineering of a variety of tissues such as cartilage, bone, tendon, ligament, skin, blood vessels, nerve, bladder, tooth, etc. This chapter reviews some recent development of collagen-based scaffolds and their application in tissue engineering.

15.2

Structure and properties of collagen

Collagen is the most abundant and ubiquitous protein in vertebrates.1 It accounts for 20-30% of total body proteins and its functions range from serving crucial biomechanical purposes in skin, bone, cartilage, tendon, and ligament to controlling cellular gene expression in development. More than 20 genetically distinct collagens have been identified.2–4 Type I collagen occurs throughout the body, except in cartilage. It is the principal collagen in the dermis, fasciae and tendons and is a major component of mature scar tissue. Type II collagen occurs in cartilage, the developing cornea, and in the vitreous body of the eye. Type III collagen dominates in the walls of blood 396 © 2008, Woodhead Publishing Limited

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vessels and hollow intestinal organs and copolymerizes with type I collagen. Types V and XI collagen are minor components and occur predominantly copolymerized with collagen I (type V) and collagen II (type XI). Collagen is comprised of three polypeptides, each having a general amino acid sequence of (-Gly-X-Y-)n, where X is any other amino acid and is frequently proline and Y is any other amino acid and is frequently hydroxyproline. Type I collagen isolated from various tissues has a molecular weight of about 283 000 Da. It is composed of three left-handed helical polypeptide chains, which are intertwined forming a right-handed helix around a central molecular axis. Two of the polypeptide chains are identical (α1) having 1056 amino acid residues; the third polypeptide chain (α2) has 1029 amino residues. Over 95% of the amino acids have the sequence Gly-X-Y. Of the remaining 5%, the molecule does not have the sequence Gly-X-Y and is therefore not triple-helical. These nonhelical portions of the molecule are located at the N- and C-terminal ends and are referred to as telopeptides (9~26 residues). The whole molecule has a length of about 280 nm and a distance of about 1.5 nm and has a conformation similar to that of a rigid rod. The telopeptides are regions where intermolecular crosslinks are formed in vivo. The method commonly used to solubilize the collagen molecules from crosslinked fibrils with proteolytic enzymes such as pepsin removes the telopeptides from the collagen molecule. Pepsin-solubilized collagen is referred to as atelocollagen. Collagen has been widely used to prepare scaffolds for tissue engineering.5–8 It can have the form and shape of natural tissue, where other molecular and cellular components have been removed by biochemical treatments, while retaining the crosslinked structure of the tissue. Alternatively, it can be made as a soluble, purified product by enzymatic removal of the crosslinks of the intact tissue and solubilized in dilute acid, and then reconstituted and stabilized in the shape and form required of the scaffold. It can be prepared as porous scaffolds, gels, and hybrid scaffolds with other naturally derived polymers, biodegradable synthetic polymers, and inorganic biomaterials.

15.3

Collagen sponge

Collagen sponges are generally formed by freeze-drying an aqueous collagen solution.9–11 The freeze-drying process includes freezing an aqueous solution of collagen or collagen gel at a low temperature and subsequent sublimation of the ice crystals by vacuum at low temperature. The freezing temperature and freezing rate have some effect on the porous structure of the resulting collagen sponge. Fast freezing at low temperatures induces cracking, uniform small channels, and the production of a fibrous structure. Slow freezing at higher temperatures results in nonuniformity and large pores with more collapsed pores than continuous channels. Unidirectionally structured collagen

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sponge has been prepared by a unidirectional freezing-drying method.12 Faraj et al. prepared three-dimensional collagen scaffolds with a specific threedimensional structural design resembling the actual extracellular matrix (ECM) of a particular tissue using specific freezing regimes.13 Collagen scaffolds resembling the cup-shaped parenchymal (alveolar) architecture of lung, scaffolds that mimic the parallel collagen organization of tendon, and scaffolds that mimic the three-dimensional organization of skin were developed. The scaffold morphology could be controlled by the freezing rate, type of suspension medium, and specific additives (e.g., ethanol). Collagen sponge scaffolds have been used for the tissue engineering of various tissues and organs. Juncosa-Melvin et al. created autogenous tissueengineered tendon constructs by seeding rabbit mesenchymal stem cells in type I collagen sponges.14 Collagen sponge has been used for the threedimensional culture of human intervertebral disc cells. Its effect was compared with other cell carriers such as collagen gel, agarose, alginate and fibrin gel.15–17 The collagen sponge and agarose were found to provide superior microenvironments for the formation of ECM. The collagen sponge provided greater cell proliferation and appeared superior to agarose. Although some investigators have been successful in using the injection of cells for disc tissue engineering,18 a collagen sponge scaffold loaded with cells facilitates the in vivo placement of the cell-carrier construct.19 A bio-artificial periosteum composed of osteogenic cells and collagen sponge was developed.20 The bio-artificial periosteum had promotion effects on in vitro and in vivo osteogenesis. Hepatic organoids were reconstructed by culturing small hepatocytes (SHs), which are hepatic progenitor cells, in a collagen sponge.21 After culture for 1 month, cell aggregates were formed in the sponge and showed characteristic tissue architecture: columnar and/or cuboidal epithelial cells lined the surface of the sponge. The cells in the collagen sponge actively proliferated and the hepatocytes excreted albumin into the medium. Sabbagh et al. used collagen sponge for the culture of urothelial cells as a preliminary step in engineering urothelial autologous grafts.22 They reported that the collagen sponges supported the growth and stratification of urothelial cells and are a suitable substrate for developing urothelial autologous grafts. Collagen sponge was used for tooth tissue engineering.23 Cells from porcine third molar teeth at the early stage of crown formation attached more quickly and their ALP activity was significantly greater for the collagen sponge scaffold than that for a polyglycolic acid fiber mesh. The result indicates that a collagen sponge scaffold allows tooth production with a higher degree of success than does polyglycolic acid fiber mesh and that collagen sponge scaffold is superior to a polyglycolic acid fiber mesh scaffold for tooth-tissue engineering. Taylor et al. cultured human cardiac valve interstitial cells (ICs) in a collagen sponge to regenerate a constructure resembling a valve leaflet.24

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Collagen sponge is a suitable biodegradable scaffold that can maintain viable valve ICs and appears to enhance the capacity of cells to express their original phenotype. Shimizu et al. used collagen sponge to regenerate tracheal tissue by employing an in situ tissue engineering technique for airway reconstruction.25–27 Based on their previous successful experimental animal studies, they applied the regenerative technique to repair the trachea of a 78year-old woman with thyroid cancer. A Marlex mesh tube covered with collagen sponge was used as a tissue scaffold. Good epithelialization has been observed on the tracheal luminal surface with no complications for two years. Buma et al. compared the effects of crosslinked type I and type II collagen matrices on cartilage tissue engineering.28 They concluded that different types of collagen matrices induce different tissue responses in full-thickness articular cartilage defects. Type I collagen-based matrices are superior for guiding progenitor cells from a subchondral origin into the defect. In type II collagen-based matrices, cell migration is less, but invading cells are directed into a chondrocyte phenotype. Based on these observations, it seems that a composite matrix consisting of a deep layer of type I collagen and a more superficial layer of type II collagen might be the matrix of choice for cartilage regeneration. A collagen matrix composed of two layers, namely a type I/III collagen layer and a type II collagen layer, was used to evaluate the morphologic and biochemical behavior and activity of human chondrocytes taken from nonarthritic and osteoarthritic cartilage. The type I/III collagen layer is less porous and is further divided into rough and smooth sides; the smooth side is the surface facing the articular cavity. The two collagen types can be differentiated by their different fibrillar size and electron density.29,30 The porous layer is composed of type II collagen and serves the cell-seeding process. Type II collagen has been shown to maintain the chondrocyte phenotype to a better extent than type I collagen and is therefore more suitable for cell seeding. The matrices are composed of porcine collagen. Moderate crosslinking had been achieved by ultraviolet (UV) irradiation. Chondrocytes of nonarthritic cartilage revealed a larger number of spherical cells, consistent with a chondrocytic phenotype. A biochemical assay showed a net increase in GAG content in nonarthritic chondrocytes, whereas almost no GAGs were seen in osteoarthritic cells. Human articular chondrocytes isolated from osteoarthritic cartilage seem to have less bioactivity after expansion and culture in a sponge consisting of types I, II and III collagen compared with chondrocytes from nonarthritic cartilage.31 The culture conditions and release of growth factors have been combined for culture in collagen sponge.32–34 The medium perfusion and dynamic culture conditions showed some effects on the chondrogenesis of articular chondrocytes when cultured in collagen sponges. Yates et al. assessed porous, 3D collagen sponges for in vitro engineering of cartilage under both standard and serum-free culture conditions.32 They reported that porous 3D collagen

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sponges maintain chondrocyte viability, shape, and synthetic activity by providing an environment favorable for high-density chondrogenesis and that collagen sponges have potential as scaffolds for cartilage tissue engineering. Tabata et al. combined collagen sponge with an appropriate controlled release of bFGF to achieve an in situ formation of adipose tissue in rats.35 They concluded that a combination of scaffold collagen with an appropriate controlled release of bFGF was essential for achieving the in situ formation of adipose tissue even without preadipocytes.

15.4

Collagen gel

Collagen gels are very attractive for tissue engineering applications because they can retain cells and carry bioactive molecules such as growth factors. Collagen gels can be formed by shifting the pH of dispersion away from its isoelectric point. Alternatively, the collagen material can be subjected to a chemical modification procedure to change its charge profile to a net positively charged or negatively charged protein before hydrating the material to form a gel matrix. The primary factor currently limiting the use of collagen-based gels for tissue engineering is the contraction of the gel by the seeded cells, an undesirable side effect of the structure’s poor mechanical properties.36–38 Chemical crosslinking with glutaraldehyde or diphenylphosphoryl azide can enhance the mechanical stiffness of a collagen gel, but the in vivo biocompatibility is compromised under such conditions.36 Alternatively, Gentleman et al. showed that the addition of short collagen fibers substantially decreases the amount by which fibroblasts can contract collagen gels and increases the permeability of the gels without affecting cell viability.37 Lewus et al. also demonstrated that preservation of the original dimensions can be achieved without compromising cellular viability by using short collagen fibers to create a collagen composite gel.39 Ibusuki et al. demonstrated that by photochemically crosslinking collagen gels using riboflavin and visible light, stable gel scaffolds with a favorable cell survival rate can be produced.40 Wakitani et al. developed a technique of employing collagen gels as a carrier to transplant and maintain chondrocytes in defects.41,42 They transplanted allograft articular chondrocytes embedded in collagen gel into full-thickness defects in rabbit articular cartilage to repair the cartilage. Twenty-four weeks after the transplantation, the defects were filled with hyaline cartilage, specifically synthesizing type II collagen. Ochi et al. investigated the clinical, arthroscopic, and biomechanical outcome of transplanting autologous chondrocytes, cultured in atelocollagen gel, for the treatment of full-thickness defects of cartilage in 28 knees (26 patients) over a minimum period of 25 months.43 The transplantation eliminated any locking of the knee and reduced pain and swelling in all patients. The mean Lysholm scores improved

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significantly. Arthroscopic assessment indicated that 26 knees (93%) had a good or excellent outcome. There were few adverse aspects, except for marked hypertrophy of the graft in three knees, partial detachment of the periosteum in three, and partial ossification of the graft in one. Biomechanical tests revealed that the transplants had acquired a degree of hardness similar to that of the surrounding cartilage. They concluded that transplanting chondrocytes in a newly formed matrix of atelocollagen gel can promote restoration of the articular cartilage of the knee. Yokoyama et al. reported the formation of cartilage from the composites of synovium-derived MSCs with collagen gel in vitro.44 Wakitani et al. assessed the effectiveness of autologous bone marrow stromal cell transplantation for the repair of full-thickness articular cartilage defects in the patellae of a 26-year-old female and a 44year-old male.45 Bone marrow stromal cells were embedded in a collagen gel, transplanted into the articular cartilage defect in the patellae, and covered with autologous periosteum. Six months after transplantation, clinical symptoms (pain and walking ability) had improved significantly and the improvement has remained in effect (five years and nine months posttransplantation in one case, and four years in the other). Both patients are satisfied with the outcome. As early as two months after transplantation, the defects were covered with tissue that showed slight metachromatic staining. Two years after the first and one year after the second transplantation, arthroscopy was performed and the defects were repaired with fibrocartilage. The results indicate that autologous bone marrow stromal cell transplantation is an effective approach in promoting the repair of articular cartilage defects. Wiesmann used collagen gel for bone tissue engineering. Periosteal-derived osteoblasts were cultured for up to three weeks in a three-dimensional collagen gel.46 Osteoblasts proliferated in the collagen gel without loss of viability during the entire experimental period. They demonstrated a mature osteoblast phenotype as indicated by the synthesis of a bone-like extracellular matrix. They formed an extracellular matrix containing osteocalcin, osteonectin, and newly synthesized type I collagen. Saadeh et al. applied type I collagen gel in repairing critical-sized mandibular defects in rats.47 They demonstrated the ability of type I collagen gel to promote the healing of a membranous bony defect that would not otherwise be healed at six weeks. The suitability of type I collagen as a carrier matrix provides many opportunities for tissueengineered approaches to further facilitate the healing of bony defects. Promoting bone formation through tissue engineering matrices offers great promise for skeletal healing and reconstruction. Collagen-populated hydrated gels have been used in the treatment of burn patients or chronic wounds.48 The skin substitutes are produced by culturing keratinocytes on a matured dermal equivalent composed of fibroblasts included in a collagen gel. Collagen gels have also been used for the tissue engineering of heart valves49 and anterior cruciate ligaments.50

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The mechanical properties of collagen gels may affect the formation of three-dimensional networks of tissues or organs. Bovine pulmonary microvascular ECs (BPMECs) were cultured on a series of collagen gels of different degrees of stiffness but the same collagen concentration.51 Imaging techniques revealed that cells cultured in rigid and flexible gels formed 3D networks via different processes: cells formed dense, thin networks in the flexible gels, whereas they formed thicker and deeper networks in the rigid gels. Cross-sections of the networks revealed that those formed within the rigid gels had large lumens composed of multiple cells, whereas those formed within the flexible gels had small, intracellular vacuoles. Seliktar et al. reported that dynamic mechanical conditioning improved the mechanical properties of tissue-engineered blood vessel constructs composed of living cells embedded in a collagen-gel scaffold.52 Dynamic mechanical conditioning during tissue culture led to an improvement in the properties of tissueengineered blood vessel constructs in terms of mechanical strength and histological organization.

15.5

Collagen–glycosaminoglycan (GAG) scaffolds

Collagen–GAG scaffolds have been used extensively for tissue engineering because they can be manufactured with a variety of pore structures and a large range of degradation rates and can be sterilized using heat or chemical procedures.53 The collagen–GAG scaffold is fabricated using a freeze-drying process in which a collagen–GAG suspension is frozen, leaving the collagen– GAG to coprecipitate between growing ice crystals.54 The conventional freezedrying process for fabricating collagen–glycosaminoglycan scaffolds creates variable cooling rates throughout the scaffold during freezing, producing a heterogeneous matrix pore structure with large variations in average pore diameters at different locations throughout the scaffold. Collagenglycosaminoglycan scaffolds with a homogeneous structure were prepared using a modified method by generating more homogeneous freezing by controlling the freezing rate and obtaining more uniform contact between the pan containing the collagen–glycosaminoglycan suspension and the freezing shelf with smaller, less warped pans. Compared with the pores prepared in scaffolds produced using conventional freezing, the pores produced using the new technique appeared to be more equiaxed. An electrospinning technique can be used to fabricate nanofibrous collagen– GAG scaffolds.55 Such scaffolds were obtained by electrospinning collagen blended with chondroitin sulfate (CS), a widely used GAG, in a mixed solvent of trifluoroethanol and water. The electrospun collagen–GAG scaffold with 4% CS (COLL-CS-04) exhibited a uniform fiber structure with nanoscale diameters. A second collagen–GAG scaffold with 10% CS consisted of smaller diameter fibers but exhibited a broader diameter distribution than that of

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COLL-CS-04 due to the different solution properties. After crosslinking with glutaraldehyde vapor, the collagen–GAG scaffolds became more biostable and were resistant to collagenase degradation. The potential of applying nanoscale collagen–GAG scaffolds in tissue engineering is significant since the nanodimensional fibers made of natural ECM closely mimic the native ECM found in the human body. The high surface area characteristic of this scaffold may maximize cell-ECM interaction and promote tissue regeneration faster than that of other conventional scaffolds. The pore size of collagenglycosaminoglycan has been reported to affect cell functions. It has been shown that collagen–glycosaminoglycan scaffolds used for studies of skin regeneration were inactive when the mean pore size was either lower than 20 µm or higher than 120 µm.56 O’Brien reported on the relationship between cell attachment and viability in scaffolds and the pore structure of collagen– glycosaminoglycan scaffolds.57 Collagen–glycosaminoglycan scaffolds with a constant composition and solid volume fraction, but with four different pore sizes corresponding to four levels of specific surface area, were manufactured using a freeze-drying technique and used for culture of MC3T3E1 mouse clonal osteogenic cells. Cell attachment and viability were primarily influenced by scaffold surface area over the pore size range of 95.9–150.5 µm for MC3T3 cells. The incorporation of GAG into a collagen scaffold has been found to improve tissue growth and regeneration over the use of collagen alone. However, the role of the GAG component, most frequently chondroitin-6sulfate, in collagen–GAG scaffold is not completely understood. Although GAGs are recognized to be water-binding and indirectly participate in collagen fibril organization, the presence of GAGs has not been found to alter the morphology, in vitro degradation, or mean tensile strength of native unseeded collagen–GAG matrices.58,59 The GAG component, however, has been observed to delay matrix degradation in vivo, which indicates some possible interaction between cells and the GAGs. The collagen–GAG combination also induces more native tissue-like ECM composition in the engineered tissue than would occur in collagen alone, possibly through interactions with growth factors. Furthermore, the collagen–GAG scaffolds cause cells to retain more proteoglycan aggregates within the scaffold.60 Collagen–GAG scaffolds have been used as scaffolds for osteogenic, chondrogenic, and lung tissue development. Rat MSCs were cultured in a novel collagen–glycosaminoglycan scaffold in the presence of a standard combination of osteoinductive factors.61 The initial response of the cells in 3D collagen seemed to be faster than cells cultured in 2D collagen, as evidenced by collagen type I expression. Later markers showed that the osteogenic differentiation of MSCs took longer in the 3D environment of a collagen GAG scaffold compared to that in standard 2D culture conditions. Furthermore, it was shown that complete scaffold mineralization could be evoked within

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a six-week period. The MSC-seeded collagen GAG scaffolds can be used for bone tissue engineering applications. Vickers et al. prepared collagen–GAG scaffolds by chemical crosslinking to achieve a range of crosslink densities.62 Chondrocyte-seeded scaffolds of varying crosslink densities were then cultured for two weeks to evaluate the effects of crosslink density on scaffold contraction and chondrogenesis. Scaffolds with low crosslink densities experienced cellmediated contraction, increased cell number densities, a greater degree of chondrogenesis, and an apparent increase in the rate of degradation of the scaffold compared to more highly crosslinked scaffolds, which resisted cellular contraction. The results suggest the promise of ‘dynamic pore reduction’ for scaffolds used in articular cartilage tissue engineering. In this approach, scaffolds would have an initial pore diameter large enough to facilitate cell seeding and a mechanical stiffness low enough to allow cell-mediated contraction to yield a reduced pore volume that favors chondrogenesis. This approach might provide a useful alternative to traditional means of increasing cell numbers and cell density and retention of synthesized molecules that promote cartilage formation in tissue-engineered constructs. Chen et al. used collagen–GAG scaffold to culture fetal rat lung cells.63 All of the cell-seeded scaffolds underwent cell-mediated contraction that appeared to be associated with the expression of alpha-smooth muscle actin in some cells. These results demonstrated the capability of dissociated lung cells to form lung histotypic structures in collagen–GAG tissue-engineering scaffolds in vitro. This culture system might be of value in facilitating exploration of strategies for preparing such scaffolds for the regeneration of lung tissue in vivo.

15.6

Acellularized scaffolds

Acellularized collagen scaffolds are prepared by decellularizing tissues and organs while retaining the extracellular matrices.64 Removal of cells from a tissue or organ leaves a complex mixture of structural and functional proteins that constitute the extracellular matrix (ECM). The collagen is in a complex network with other ECM molecules. The tissues from which the ECM is harvested, the species of origin, the decellularization methods, and the methods of terminal sterilization for the acellularized scaffolds vary widely. The efficiency of cell removal from a tissue is dependent on the origin of the tissue and the specific physical, chemical and enzymatic methods that are used. Each of these treatments affects the biochemical composition, tissue ultrastructure, and mechanical behavior of the remaining extracellular matrix (ECM) scaffold which, in turn, affect the host response to the material. Acellularized scaffolds derived from decellularized tissues and organs have been used successfully in both pre-clinical animal studies and in human clinical applications. ECM from a variety of tissues, including heart valves, blood vessels, skin, nerves, skeletal muscle, tendon, ligament, small intestinal

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submucosa, urinary bladder, vocal fold and liver have been studied for tissue engineering and regenerative medicine applications.

15.7

Hybrid scaffolds

Scaffolds constructed from collagen and hybrid scaffolds constructed with other naturally derived polymers such as GAG have poor mechanical strength. It is difficult to maintain the new tissue in the initially designed shape because these scaffolds are too weak to withstand cell contraction during tissue regeneration. The low mechanical strength of these scaffolds also makes medical manipulation difficult and unable to keep the space necessary for new in vivo tissue formation. To solve these problems, collagens have been hybridized with mechanically strong materials into hybrid structures.65 One such major set of scaffolds is the hybrid scaffolds of collagen with biodegradable synthetic polymers such as polyesters and polycaprolactones. Polyesters such as poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and their copolymers of poly(lactic-co-glycolic acid) (PLGA) are most commonly used for tissue engineering and hybridization with collagen. Chen et al. developed a novel hybridization method by forming collagen microsponges in the openings of a synthetic polymer skeleton (Fig. 15.1). The synthetic polymer skeleton enables easy formation into the desired shapes and provides the appropriate mechanical strength, while the nested microsponges of naturally derived polymers facilitate cell seeding and cell attachment.66 Several kinds of such hybrid scaffolds were developed (Fig. 15.2). Synthetic polymer sponge

Microsponge of naturally derived polymer

Hybrid sponge Introduction of naturally derived polymer microsponge in the pores

Formation of naturally derived polymer microsponge in the openings

Synthetic polymer mesh

Hybrid mesh

15.1 Hybridization of biodegradable synthetic polymers and naturally derived polymers

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(a)

(b)

(c)

(d)

15.2 SEM photomicrographs of (a) PLGA-collagen hybrid sponge; (b) PLGA-collagen hybrid mesh; (c) PLLA-collagen hybrid braid; and (d) collagen/PLGA-collagen biphasic sponge.

One example is a hybrid sponge prepared by introducing collagen microsponges in the pores of a PLGA sponge.67 The PLGA-collagen hybrid sponge was prepared by immersing a PLGA sponge in a bovine collagen type I acidic solution under negative pressure, freezing at –80°C, and freezedrying. The hybrid sponge was further crosslinked by treating with glutaraldehyde vapor and washing with glycine aqueous solution and water. The hybrid structure of the PLGA-collagen hybrid sponge was confirmed by scanning electron microscopy. Collagen microsponges with interconnected pore structures were formed in the pores of the PLGA sponge (Fig. 15.2a). SEM-electron probe microanalysis of elemental nitrogen confirmed that the collagen microsponges were formed in the pores of the PLGA sponge and that the pore surfaces were coated with collagen. The ultimate tensile strength, the modulus of elasticity, and the static stiffness of the PLGA-collagen hybrid sponge were higher than were those of PLGA and collagen sponges, in both dry and wet states. The second example is a PLGA-collagen hybrid mesh.68 This was prepared by forming collagen microsponges in the openings of a knitted PLGA mesh. SEM observation confirmed that web-like collagen

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microsponges were formed in the openings of the synthetic PLGA mesh (Fig. 15.2b). The moduli of elasticity of the hybrid mesh, PLGA mesh, and collagen sponge were 35.4 ±1.4, 35.2 ±1.0 and 0.020 ±0.001 MPa, respectively. The hybrid mesh exhibited a significantly higher tensile strength than did the collagen sponge alone, and was similar to the PLGA mesh. The third example is a hybrid braid formed by introducing collagen microsponges in the interstices of a poly(L-lactic acid) (PLLA) braid (Fig. 15.2c).69 The fourth example is a PLGA/collagen biphasic scaffold (Fig. 15.2d).70 The biphasic scaffold is composed of an upper layer of collagen sponge and a lower layer of PLGAcollagen hybrid sponge. It was developed as follows. First, a biodegradable PLGA sponge cylinder was prepared by adding NaCl particulates to a PLGA solution in chloroform and leaching them out of the dried PLGA/NaCl composite. Then, a collagen/PLGA-collagen biphasic sponge cylinder was prepared by introducing collagen sponge into the pores of the PLGA sponge and forming a collagen sponge at one side of the PLGA sponge. These hybrid scaffolds have been used for tissue engineering of skin,71 cartilage,72–74 bone,75,76 osteochondral tissue,77 ligament,78 bladder,79 and cardiovascular tissue.80 Chen et al. reported cartilage tissue engineering using PLGA-collagen hybrid mesh.73 The PLGA-collagen hybrid mesh was used for three-dimensional culture of bovine articular chondrocytes. Subcultured bovine articular chondrocytes were seeded into the PLGA-collagen hybrid mesh and cultured in vitro in culture media in a 5% CO2 atmosphere at 37°C. The chondrocytes adhered to the hybrid mesh, proliferated, and regenerated a cartilaginous matrix filling the voids in the hybrid mesh. The web-like collagen microsponge that formed in the openings of the knitted mesh not only prevented the seeded cells from going through the composite web, but also increased the specific surface area to provide sufficient surfaces for spatially even chondrocyte distribution. After being cultured in vitro for 1 day, the cell/scaffold sheets were used singly to regenerate thin cartilage implants or in laminated form to yield thick cartilage implants. The thickness of the implant could be controlled by changing the number of sheets of laminated scaffold. The cell/scaffold sheets could also be rolled up in the shape of a cylinder in which case the thickness of the implant was adjusted by the height of the roll and its diameter by the number of sheets in the roll. Round, disk-shaped single-sheet, five-sheet, and 8 mm-high roll implants were cultured in DMEM for another week and implanted subcutaneously in the dorsum of athymic nude mice. The implants were harvested after 4, 8, and 12 weeks. Gross examination of these grafts showed that the implants retained their original shapes for all implantation periods and appeared pearly white (Fig. 15.3). The thickness of the engineered cartilage implants of single-sheet, five-sheet, and rolled implants were 200 µm, 1 mm and 8 mm, respectively. Histological examination of these specimens using hematoxylin and eosin stains indicated a uniform spatial cell distribution throughout all

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15.3 Gross appearance of tissue-engineered cartilage.

the implants, both radially and longitudinally. The laminated and rolled implant sheets became integrated with each other. All the chondrocytes in the implants remained viable, proliferating and secreting extracellular matrix components to form homogeneously compact cartilage tissues. In all the implants, the chondrocytes showed a natural round morphology. A bright safranin-O-positive stain indicated that glycosaminoglycans (GAG) were abundant and homogeneously distributed throughout the implants. Toluidine blue staining demonstrated the typical metachromasia of articular cartilage, coinciding with the results of the safranin-O staining. Immunohistological staining with an antibody to type II collagen showed a homogeneous extracellular staining for type II collagen. The similarity of the results for the five-sheet and roll implants to those of the single-sheet implant suggests that an increase in the implant thickness from 200 µm to 8 mm does not compromise cell viability, cell uniformity, or cellular function. The mechanical properties of the fivesheet implant after 12 weeks and bovine native articular cartilage were evaluated by a dynamic compression mechanical test using a viscoelastic spectrometer. The dynamic complex modulus (E*), structural stiffness, and phase lag (tanδ) measured at 11 Hz reached 37.8%, 57.0% and 86.3% of those of native bovine articular cartilage, respectively. These results suggest the formation of articular cartilage. The spatially even distribution of a sufficient number of chondrocytes facilitated the regeneration of articular cartilage. Articular cartilage patches with a thickness ranging from 200 µm to 8 mm were produced by laminating or rolling the hybrid mesh sheets.

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Tabata et al. developed a collagen sponge reinforced by the incorporation of poly(glycolic acid) (PGA) fiber.81–83 The hybrid scaffold was used for the culture of rat MSCs for bone tissue engineering. A hybrid scaffold of poly(DL-lactide-co-caprolactone), poly(DL-lactideco-glycolide) (PLGA), and type I collagen with open interconnected pores and an average void volume of 80 ±5% was developed and used for rat cardiac tissue engineering.84 Construct cellularity, the presence of cardiac markers, and contractile properties were markedly improved in the composite scaffolds as compared with those in both collagen sponge and PLGA sponge. Polyglycolic acid (PGA)-collagen hybrid tubes have been developed for nerve tissue engineering.85–86 Not only collagen sponge, but also collagen gel has been hybridized with biodegradable synthetic polymers to improve the mechanical properties.87–89 Collagen has also been hybridized with other polymers such as chitosan and used for cartilage and skin tissue engineering, and polypropylene.90–93

15.8

Future trends

Collagen is one of the most widely used naturally derived polymers for manufacturing tissue scaffolds. It is a primary structural protein of native ECM. It has a variety of functional properties favorable for cell adhesion, proliferation, differentiation and ECM secretion. Collagen-based scaffolds have become one of the major biomaterials used for tissue engineering and have been used for the regeneration of many kinds of tissues and organs. Some of their applications are very exciting and have already been used for clinical treatments, but some are still in the preliminary stages. Challenges remain for controlling the biodegradation of collagen-based scaffolds and improving their mechanical properties. The contraction and deformation of collagen-based scaffolds has limited their application to load-bearing tissues. Another problem concerning collagen-based scaffolds is their origin. The most frequently used collagen is from porcine or bovine sources, or other similar animals. To avoid problems associated with the possible transfer of pathogens, safer sources such as human recombinant collagen or chemically synthesized collagen are preferred.

15.9

References

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38 Liu X, Umino T, Cano M, Ertl R, Veys T, Spurzem J, Romberger D and Rennard S I (1998), Human bronchial epithelial cells can contract type I collagen gels, Am J Physiol, 274, 58–65. 39 Lewus K E and Nauman E A (2005), In vitro characterization of a bone marrow stem cell-seeded collagen gel composite for soft tissue grafts, effects of fiber number and serum concentration, Tissue Eng, 11(7-8), 1015–1022. 40 Ibusuki S, Halbesma G J, Randolph M A, Redmond R W, Kochevar I E and Gill T J (2007), Photochemically cross-linked collagen gels as three-dimensional scaffolds for tissue engineering, Tissue Engineering, 13(8), 1995–2001. 41 Wakitani S, Kimura T, Hirooka A, Ochi T, Yoneda M, Yasui N, Owaki H and Ono K (1989), Repair of rabbit articular surfaces with allograft chondrocytes embedded in collagen gel, J Bone Joint Surg Br, 71(1), 74–80. 42 Kawamura S, Wakitani S, Kimura T, Maeda A, Caplan A I, Shino K and Ochi T (1998), Articular cartilage repair. Rabbit experiments with a collagen gel-biomatrix and chondrocytes cultured in it, Acta Orthop Scand, 69(1), 56–62. 43 Ochi M, Uchio Y, Kawasaki K, Wakitani S and Iwasa J (2002), Transplantation of cartilage-like tissue made by tissue engineering in the treatment of cartilage defects of the knee, J Bone Joint Surg Br, 84(4), 571–578. 44 Yokoyama A, Sekiya I, Miyazaki K, Ichinose S, Hata Y and Muneta T (2005), In vitro cartilage formation of composites of synovium-derived mesenchymal stem cells with collagen gel, Cell Tissue Res, 322(2), 289–298. 45 Wakitani S, Mitsuoka T, Nakamura N, Toritsuka Y, Nakamura Y and Horibe S (2004), Autologous bone marrow stromal cell transplantation for repair of fullthickness articular cartilage defects in human patellae, two case reports, Cell Transplant, 13(5), 595–600. 46 Wiesmann H P, Nazer N, Klatt C, Szuwart T and Meyer U (2003), Bone tissue engineering by primary osteoblast-like cells in a monolayer system and 3-dimensional collagen gel, J Oral Maxillofac Surg, 61(12), 1455–1462. 47 Saadeh P B, Khosla R K, Mehrara B J, Steinbrech D S, McCormick S A, DeVore D P and Longaker M T (2001), Repair of a critical size defect in the rat mandible using allogenic type I collagen, J Craniofac Surg, 12(6), 573–579. 48 Auger F A, Rouabhia M, Goulet F, Berthod F, Moulin V and Germain L (1998), Tissue-engineered human skin substitutes developed from collagen-populated hydrated gels, clinical and fundamental applications, Med Biol Eng Comput, 36(6), 801–812. 49 Seliktar D, Black R A, Vito R P and Nerem R M (2000), Dynamic mechanical conditioning of collagen-gel blood vessel constructs induces remodeling in vitro, Ann Biomed Eng, 28(4), 351–362. 50 Nöth U, Schupp K, Heymer A, Kall S, Jakob F, Schütze N, Baumann B, Barthel T, Eulert J and Hendrich C (2005), Anterior cruciate ligament constructs fabricated from human mesenchymal stem cells in a collagen type I hydrogel, Cytotherapy, 7(5), 447–545. 51 Yamamura N, Sudo R, Ikeda M and Tanishita K (2007), Effects of the mechanical properties of collagen gel on the in vitro formation of microvessel networks by endothelial cells, Tissue Eng, 13(7), 1443–1453. 52 Seliktar D, Black R A, Vito R P and Nerem R M (2000), Dynamic mechanical conditioning of collagen-gel blood vessel constructs induces remodeling in vitro, Ann Biomed Eng, 28(4), 351–362. 53 O’Brien F J, Harley B A, Yannas I V and Gibson L J (2005), The effect of pore size on cell adhesion in collagen–GAG scaffolds, Biomaterials, 26, 433–441.

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54 O’Brien F J, Harley B A, Yannas I V and Gibson L (2004), Influence of freezing rate on pore structure in freeze-dried collagen–GAG scaffolds, Biomaterials, 25(6), 1077– 1086. 55 Zhong S, Teo W E, Zhu X, Beuerman R, Ramakrishna S and Yung L Y (2005), Formation of collagen–glycosaminoglycan blended nanofibrous scaffolds and their biological properties, Biomacromolecules, 6(6), 2998–3004. 56 Yannas I V, Lee E, Orgill D P, Skrabut E M and Murphy G F (1989), Synthesis and characterization of a model extracellular matrix that induces partial regeneration of adult mammalian skin, Proc Natl Acad Sci USA, 86(3), 933–937. 57 O’Brien F J, Harley B A, Yannas I V and Gibson L J ( ), The effect of pore size on cell adhesion in collagen–GAG scaffolds, Biomaterials, 26(4), 433–441. 58 Pieper J S, Oosterhof A, Dijkstra P J, Veerkamp J H and Van Kuppevelt T H (1999), Preparation and characterization of porous crosslinked collagenous matrices containing bioavailable chondroitin sulphate, Biomaterials, 20, 847–858. 59 Daamen W F, Van Moerkerk H T, Hafmans T, Buttafoco L, Poot A A, Veerkamp J H and Van Kuppevelt T H (2003), Preparation and evaluation of molecularly-defined collagenelastin-glycosaminoglycan scaffolds for tissue engineering, Biomaterials, 24, 4001–4009. 60 Rong Y, Sugumaran G, Silbert J E and Spector M (2002), Proteoglycans synthesized by canine intervertebral disc cells grown in a type I collagen–glycosaminoglycan matrix, Tissue Eng, 8, 1037–1047. 61 Farrell E, Byrne E M, Fischer J, O’Brien F J, O’Connell B C, Prendergast P J and Campbell V A (2007), A comparison of the osteogenic potential of adult rat mesenchymal stem cells cultured in 2-D and on 3-D collagen glycosaminoglycan scaffolds, Technol Health Care, 15(1), 19–31. 62 Vickers S M, Squitieri L S and Spector M (2006), Effects of cross-linking type II collagen–GAG scaffolds on chondrogenesis in vitro: dynamic pore reduction promotes cartilage formation, Tissue Eng, 12(5), 1345–1355. 63 Chen P, Marsilio E, Goldstein R H, Yannas I V and Spector M (2005), Formation of lung alveolar-like structures in collagen–glycosaminoglycan scaffolds in vitro, Tissue Eng, 11(9-10), 1436–1448. 64 Gilbert T W, Sellaro T L and Badylak S F (2006), Decellularization of tissues and organs, Biomaterials, 27(19), 3675–3683. 65 Chen G, Ushida T and Tateishi T (2002), Scaffold design for tissue engineering, Macromolecular Bioscience, 2(1), 67–77. 66 Chen G, Ushida T and Tateishi T (2000), Hybrid biomaterials for tissue engineering: A Preparative Method of PLA or PLGA-Collagen Hybrid Sponge, Advanced Materials, 12(6), 455–457. 67 Chen G, Ushida T and Tateishi T (2000), A biodegradable hybrid sponge nested with collagen microsponges, Journal of Biomedical Materials Research, 51(2), 273–279. 68 Chen G, Ushida T and Tateishi T (2000), A hybrid network of synthetic polymer mesh and collagen sponge, Chemical Communications, 16, 1505–1506. 69 Ide A, Sakane M, Chen G, Shimojo H, Ushida T, Tateishi T, Wadano Y and Miyanaga Y (2001), Collagen hybridization with the poly-L lactic acid (PLLA) braid promotes ligament cell migration, Mater Sci & Eng C, 17(1-2), 95–99. 70 Chen G, Sato T, Tanaka J and Tateishi T (2006), Preparation of a biphasic scaffold for osteochondral tissue engineering, Materials Science & Engineering C-Biomimetic and Supramolecular Systems, 26(1), 118–123. 71 Chen G, Sato T, Hajime H, Ushida T, Tateishi T and Tanaka T (2005), Culturing of

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85 Ito T, Nakamura T, Takagi T, Toba T, Hagiwara A, Yamagishi H and Shimizu Y (2003), Biodegradation of polyglycolic acid-collagen composite tubes for nerve guide in the peritoneal cavity, ASAIO J, 49(4), 417–421. 86 Ito T, Nakamura T, Suzuki K, Takagi T, Toba T, Hagiwara A, Kihara K, Miki T, Yamagishi H and Shimizu Y (2003), Regeneration of hypogastric nerve using a polyglycolic acid (PGA)-collagen nerve conduit filled with collagen sponge proved electrophysiologically in a canine model, Int J Artif Organs, 26(3), 245–251. 87 Ushida T, Furukawa K, Toita K and Tateishi T (2002), Three-dimensional seeding of chondrocytes encapsulated in collagen gel into PLLA scaffolds, Cell Transplant, 11(5), 489–494. 88 Hannouche D, Terai H, Fuchs J R, Terada S, Zand S, Nasseri B A, Petite H, Sedel L and Vacanti J P (2007), Engineering of implantable cartilaginous structures from bone marrow-derived mesenchymal stem cells, Tissue Eng, 213(1), 87–99. 89 Ito Y, Ochi M, Adachi N, Sugawara K, Yanada S, Ikada Y and Ronakorn P (2005), Repair of osteochondral defect with tissue-engineered chondral plug in a rabbit model, Arthroscopy, 21(10), 1155–1163. 90 Shi D H, Cai D Z, Zhou C R, Rong L M, Wang K and Xu Y C (2005), Development and potential of a biomimetic chitosan/type II collagen scaffold for cartilage tissue engineering, Chin Med J (Engl), 118(17), 1436–1443. 91 Gingras M, Paradis I and Berthod F (2003), Nerve regeneration in a collagenchitosan tissue-engineered skin transplanted on nude mice, Biomaterials, 24(9), 1653–1661. 92 Shi D H, Cai D Z, Zhou C R, Rong L M, Wang K and Xu Y C (2005), Development and potential of a biomimetic chitosan/type II collagen scaffold for cartilage tissue engineering, Chin Med J (Engl), 118(17), 1436–1443. 93 Nakashima S, Nakamura T, Miyagawa K, Yoshikawa T, Kin S, Kuriu Y, Nakase Y, Sakakura C, Otsuji E, Hagiwara A and Yamagishi H (2007), In situ tissue engineering of the bile duct using polypropylene mesh-collagen tubes, Int J Artif Organs, 30(1), 75–85.

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16 Polyhydroxyalkanoate and its potential for biomedical applications P. F U R R E R and M. Z I N N, Swiss Federal Laboratories for Materials Testing and Research (Empa), Switzerland, and S. P A N K E, Swiss Federal Institute of Technology (ETH), Switzerland

16.1

Introduction

Poly([R]-hydroxyalkanoate) (PHA) is a water insoluble, biodegradable, and biocompatible polyester that is accumulated by a large number of bacteria as a carbon storage compound under nutrient limited growth conditions. PHA has a wide variety of potential applications in medicine and pharmacy due to its biodegradability, biocompatibility and broad spectrum of physical properties. For this purpose PHA of high purity is needed and adequate methods for the recovery and purification are crucial. This chapter describes the biosynthesis, the fermentative production and recovery, and the potential medical applications of PHA. New biomaterials of the third generation are needed for particular medical applications such as vascular implants, heart valves and cardiovascular fabrics.1–3 They are designed to stimulate specific cellular responses at the molecular level and to combine the concepts of bioactive and resorbable materials. They are supposed to help the body heal itself by prompting cells to repair their own tissues. Today’s polymeric and biodegradable systems used in medicine are mainly based on poly(lactic acid) (PLA), on poly(glycolic acid) (PGA), and on their co-polymers.4 Other biodegradable polymers have been proposed, but could not enter the market yet, due to lacking FDA approval.4 One of the candidates is poly([R]-hydroxyalkanoate) (PHA), a class of biodegradable and biocompatible polyesters with many potential applications in the medical field, such as heart valve scaffolds,5,6 pulmonary conduits,7 sutures, screws, bone plates, repair patches, stents, bone marrow scaffolds, and many others over the last years, as recently reviewed by Chen and Wu.8 PHA is composed of 3-, 4-, or rarely 5-hydroxy fatty acid monomers, which form linear polyesters. The general structure of poly([R]-hydroxyalkanoate) is shown in Fig. 16.1. PHAs can be separated into three classes according to the size of comprising monomers: short-chain-length PHAs (scl-PHA) with monomers of 3-5 carbon atoms, medium-chain-length PHAs (mcl-PHA) 416 © 2008, Woodhead Publishing Limited

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O R

O m

n

m = 1–3

16.1 General structure of poly ([R]-hydroxyalkanoate).

with monomers of 6-14 carbon atoms and long-chain-length PHAs (lclPHAs) with monomers of more than 14 carbon atoms. PHA is produced as reserve material by many archeae and eubacteria in aerobic and anaerobic habitats. PHA accumulation occurs when microorganisms experience a metabolic stress, such as limitation by an essential nutrient in an excess of a suitable carbon source.9 The purity of PHA is a crucial factor for sophisticated applications. It is determined by the fermentation process and the downstream processing. The production strain and the selective recovery procedure are essential factors for obtaining PHA of high quality. Although there has been a continuous progress in downstream processing of PHA for bulk applications (more than 50 patents have been filed in the past 40 years) further improvements for high-quality PHAs and in particular for mcl-PHAs are to be expected. The restricted availability of PHAs has been limiting research significantly. Poly(3-hydroxybutyrate) (PHB) and poly(3-hydroxybutyrate-co-3hydroxyvalerate) (PHBV) are the only PHA-polymers that are currently commercially available (e.g. from Metabolix (USA), Biocycle (UK), Tiannan (CN), and Biomer (D)). At present, polymers based on poly(4-hydroxybutyrate) (P4HB) are developed for medical use (www.tepha.com) and one of them has been approved by FDA recently.10 In 2008, Telles, a joint venture between Metabolix and Archer Daniels Midland, plans to start the production of PHA under the name MirelTM at an annual rate of 50 kt. PHAs offer a wide range of physical properties due to their chemical diversity and due to chemical modifications of functional groups following biosynthesis.

16.2

Biosynthesis

PHAs are a class of polyesters produced as reserve materials by many archae and eubacteria (Gram-negative and Gram-positive, see Table 16.1) in aerobic and anaerobic habitats. Up to date, more than 300 microorganisms are known to synthesize PHA.11 PHA accumulation is generally triggered when the microorganisms experience a metabolic stress, such as limitation by an essential

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Table 16.1 Important PHA-producing genera, substrates and resulting PHAs Genus

Gram stain

Substrates

PHA

Ref.

Alcaligenes Azospirillum Bacillus

– – +

Fatty acids Fatty acids, saccharides Fatty acids (C2-C6), saccharides, lactones Glucose Saccharides Saccharides Saccharides 1,3-propandiol Acetate Saccharides Saccharides, fatty acids

PHB PHB PHB, PHBV, PHBHx, P4HB PHB PHB PHB, PH4PE PHB Scl- and mcl-PHA PHB PHBV PHB, PHBV, P4HB

22, 23 24 25–29 30 31 32 33 34 35 36 37–39

Saccharides Alcohols, fatty acids Fatty acids, saccharides Fatty acids, saccharides Saccharides Fatty acids Saccharides, fatty acids Saccharides Saccharides Saccharides

PHB, PHBV PHB, PHBV PHB, PHBV PHHp-PHDd PHBV PHB, PHBV PHBV, PHPi PHBV PHB PHB

40, 41 42 43 44 45 46 47, 48 49 50 51

Beijerinckia – Brevundimonas – Burkholderia – Caulobacter – Chromobacterium – Clostridium + Corynebacterium + Cupriavidus – (former Wautersia) Haloferax – Methylobacterium – Nocardia + Pseudomonas – Rhizobium – Rhodobacter – Rhodococcus + Sphingopyxis – Staphylococcus + Streptomyces +

Notes: PHB: Poly(3-hydroxybutyrate), PHBV: Poly(3-hydroxybutyrate-co-3hydroxyvalerate), PHPi: Poly(3-hydroxypivalate), P4HB: Poly(4-hydroxybutyrate), PHBHx: Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate), P3HB-4HB: Poly(3hydroxybutyrate-co-4-hydroxybutyrate), PHHp: Poly(3-hydroxyheptenoate), PHDd: Poly(3-hydroxydodecanoate), PH4PE: Poly(3-hydroxypent-4-enoate)

nutrient (e.g., nitrogen, phosphorus or oxygen) and a concomitant excess of a suitable carbon source.12 PHA is stored intracellularly in the form of granules as a carbon and energy storage compound. The PHA granules are surrounded by a phospholipid monolayer and several proteins such as PHA polymerase and depolymerase and structural proteins called phasins (see Fig. 16.2). Upon carbon starvation or a change of the environmental pH,13 intracellular PHA depolymerases, which are attached to the granule, release 3-hydroxyalkanoic acids. PHAs have the general chemical structure depicted in Fig. 16.1. The high stereoselectivity of the producing enzyme machinery guarantees complete stereospecificity (all chiral carbon atoms in the backbone are in the R configuration), which is essential for the biodegradability and biocompatibility of PHA.14,15 The type of bacterium and the growth conditions determine the chemical composition of PHAs and the molecular weight, which typically ranges from 2 × 105 to 3 × 106 Da.16

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PHA Phasins/proteins PHA polymerase PHA depolymerase

Phospholipid monolayer

16.2 Structure of PHA- granules accumulated by bacteria131. Table 16.2 The four classes of polyester synthases (according to Rehm.17 HA-CoA = coenzyme A thioester of [R]-hydroxy fatty acids of variable length Class I

Subunits PhaC

Species

Substrate

Cupriavidus necator

3HASCL–CoA (~ C3–C5) 4HASCL–CoA 5HASCL–CoA, 3MLSCL–CoA

Pseudomonas aeruginosa

3HAMCL–CoA (~ ≥ C5)

Allochromatium vinosum

3HASCL–CoA (3HAMCL–CoA [~C6–C8] 4HA–CoA, 5HA–CoA

Bacillus megaterium

3HASCL–CoA

~ 60–73 kDa

II

PhaC

~ 60–65 kDa III

PhaC

PhaE

~ 40 kDa ~ 40 kDa IV

PhaC

PhaR

~ 40 kDa ~ 22 kDa

Research has focused on the substrate specificity of the PHA polymerase.17 Four major classes of PHA polymerases have been proposed with respect to their primary structures, their substrate specificities and their subunit composition18 (see Table 16.2). It was found that the substrate specificity of the PHA polymerase and the supply of cells with a particular substrate, control the monomeric composition of the resulting PHA. In general, different pathways for the biosynthesis of PHA are possible (see Fig. 16.3). Acetyl-Co A activated precursors of PHA are either synthesized in the course of anabolism through de novo fatty acid synthesis or stem in the course of catabolism from the β-oxidation of fatty acids that are supplied to cells. To date, more than 150 different hydroxyalkanoic acids are known as PHA constituents,18 but only few of the corresponding PHAs have been

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Carbohydrate

Fatty acid HSCoA, ATP

HSCoA

AMP + PPi, H2O O

HSCoA O

O FADH2

SCoA Acetyl-CoA CO2, ATP

O

O

1

13

SCoA R 3-ketoacyl-CoA β-oxidation

HSCoA O

O

SCoA Acetoacetyl-CoA

NADPH + H+

NADPH + H 2

+

10

SCoA R 2-trans-enoyl-CoA

OH

HSCoA R OH O O Fatty acids ACP – 19 O ACP Malonyl-ACP H2O O

9

NADP+

O

R SCoA NADPH + H (R)-3-hydroxyacyl-CoA O R

HSCoA

18

H2O

Fatty acid de novo synthesis OH

17

O

16

HSCoA

O

R ACP 3-Ketoacyl-ACP

21

HSCoA

NADPH + H+ NADP+

R ACP (R)-3-hydroxyacyl-ACP

PHA

HSCoA OH

Mcl-PHA

4–, 5–, or 6-hydroxyacyl-CoA

15

O

ACP

HSCoA

Scl-PHA

HSCoA

CO2, ACP

ACP

+

2-trans-enoyl-ACP 12

3

Other pathways

14

R

SCoA (R)-3-hydroxybutyryl-CoA

ADP + Pi

O SCoA Malonyl-CoA ACP

O

11

O

O

–

6

R SCoA (S)-3-hydroxyacyl-CoA

+

O

H2O

O

NADP+ NADP

OH

OH

7

Related carbon source

HSCoA

FAD O SCoA SCoA R 8 5 Acyl-CoA

O SCoA Acetyl-CoA

Carbohydrate

4

12

O

20

SCoA ACP R (R)-3-hydroxyacyl-CoA

16.3 Metabolic routes for PHA biosynthesis. A typical scl-PHA producer is Cupriavidus necator, whereas Pseudomonas putida GPo1 is synthesizing mcl-PHA through β-oxidation and P. putida KT2440 in addition through fatty acid de novo synthesis. Special PHA consisting of 4-, 5-, or 6-hydroxyalkanoate can be produced by various bacteria when suitable precursors are supplied. 1: βketothiolase; 2: NADPH-dependent acetoacetyl-CoA reductase; 3: scl-PHA synthase; 4: acyl-CoA synthase; 5: acyl-CoA dehydrogenase; 6: short length enoyl-CoA hydratase; 7: NAD dependent (S)-3hydroxyacyl-CoA dehydrogenase; 8: 3-ketoacyl-CoA thiolase; 9: (R)specific enoyl-CoA hydratase; 10: NADPH dependent 3-ketoacyl-CoA reductase; 11: 3-hydroxyacyl-CoA epimerase; 12: mcl-PHA synthase; 13: Acetyl-CoA carboxylase; 14: malonyl-CoA-ACP transacylase; 15: 3keto-ACP synthase; 16: 3-keto-ACP reductase; 17; 3-hydroxyacyl-ACP dehydratase; 18: enoyl-ACP reductase; 19: acyl-ACP thiolase; 20: (R)3-hydroxyacyl-ACP-CoA transacylase; 21: PHA polymerase

produced in large quantities and well characterized.19,20 As a consequence, little is known about the chemical and mechanical properties of most of these polymers. To date, PHA monomers with straight, branched, saturated and unsaturated side chains have been identified (see Fig. 16.4).21 Of special interest are functionalized groups in the side chain that allow further chemical modification, e.g. halogens, carboxy, hydroxy, epoxy, phenoxy, cyanophenoxy, nitrophenoxy, thiophenoxy, and methylester groups. The size of the monomer and its functional group considerably influence properties like the bioplastic’s melting point, the glass transition temperature, or the crystallinity, and therefore determine its final application.

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m

O

R

n

OH

HO

O

n

OH

HO

Linear and branched alkyl substituents R = alkyl substituent (C1–C12)

O OH

HO

Terminal double bond

Substituent with internal double bonds

O n

O OH

HO

O

O OH

HO

Epoxy substituent

HO

Cyclohexane substituent

OH

Aromatic substituent

X

n

HO

O

R

O OH

O

HO OH

OH

HO

R Halogen and cyano substituents X = F, Br, Cl, CN

4-Hydroxy fatty acid

5-Hydroxy fatty acid

16.4 Monomers found in poly([R]-3,4 and 5-hydroxyalkanoates).

16.2.1 Fermentative production Most PHA-producing bacteria start to accumulate PHA when their cell growth is impaired by the limitation of an essential nutrient (e.g. N, P, Mg, K, O or S) in the presence of excess carbon source. It is therefore important to use a suitable fermentation strategy to enhance the production of PHA. The PHA content is thereby the most important parameter for efficient and economic downstream processing. For fed-batch cultures a two step process is usually employed. In the first phase the cells are grown to a desired cell concentration and in the second phase PHA biosynthesis is triggered by a nutrient limitation. For continuous cultures, cell growth and PHA accumulation are controlled by the ratio of the carbon source to the limiting nutrient as well as by the dilution rate.9 PHB has been efficiently produced with Alcaligenes latus in a two-step fed-batch fermentation. Nitrogen limitation was chosen as the best strategy as it allowed the greatest enhancement of PHB production. A cell concentration of 111.7 g L–1 and a PHB content of 88 w% were reached resulting in a productivity of 4.94 g PHB L–1 h–1. 22 With Cupriavidus necator (former Wautersia eutropha

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and Ralstonia eutropha) a PHB content of 76 w% was obtained at a cell concentration of 164 g L–1 resulting in a productivity of 2.42 g L–1 h–1 under nitrogen limitation.23 PHBV has been successfully produced with C. necator. The two-step fedbatch strategy was applied using nitrogen or phosphorus limitation in the second step. The mole fraction of HV in the co-polymer could be controlled by changing the ratio of glucose to propionic acid in the feed. PHBV contents of up to 75 w% and cell concentrations of up to 158 g L–1 were obtained under nitrogen limitation.24 However, with an increasing HV mole fraction, the PHA content decreased. P3HB4HB could be produced in fed-batch cultures under nitrogen limitation. Cell concentrations of 34–49 g L–1 and PHA contents of 39–50 w% were reached with 4HB mole fractions of 1.6–25.2 mol%.25 For the efficient production of scl-PHAs, recombinant E. coli and C. necator have been intensively investigated. In contrast to natural PHAproducing bacteria, PHA accumulation by recombinant E. coli was not triggered by nutrient limitation. In a fed-batch fermentation, PHB was produced with recombinant E. coli reaching a cell concentration of 204 g L–1 with a PHB content of 77 w% and a productivity of 3.2 g L–1 h–1.26 A higher PHB productivity of 4.63 g L–1 h–1 was obtained with recombinant E. coli harboring an optimally designed plasmid containing the PHA biosynthesis genes from A. latus.27 PHBV with 11 mol% HV could be produced with a concentration of 159 g L–1 and a productivity of 2.88 g L–1 h–1 with recombinant E. coli.28 Recombinant C. necator was investigated for the production of PHB and PHBV but the final cell concentrations and PHA contents were only slightly increased compared to the parent strain.29–31 Although efforts have been made to use recombinant E. coli and recombinant Pseudomonas for the production of mcl-PHA, few fermentation processes were based on recombinant strains.32 The production of mcl-PHA was extensively investigated with Pseudomonas species.33–39 Various alkanes, alkanoic acids as well as glucose, fructose and glycerol were used as substrates for the production of mcl-PHA.40,41 Especially with structurally related carbon sources, such as alkanes and aliphatic acids, efficient mcl-PHA synthesis occurs.42 But these carbon sources are poorly water miscible or/and toxic to bacteria at rather low concentration.43 Hence the concentration of these carbon sources has to be well controlled. For fed-batch cultures the two-step fermentation strategy has often been applied. Using octanoic acid as substrate, PHA contents up to 75 w% at a cell concentration of 55 g L–1 and a productivity of 0.63 g L–1 h–1 were obtained with Pseudomonas putida GPo1 under nitrogen limitation.44 With Pseudomonas IPT 046, cell concentrations of up to 50 g L–1 with a PHA content of 63 w% and a productivity of 0.8 g L–1 h–1 were reached under phosphate limitation using a mixture of glucose and fructose as carbon source.34 When oleic acid was used as substrate for the cultivation of Pseudomonas

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putida KT2442, a cell concentration of 141 g L–1 with a PHA content of 51 w% and a productivity of 1.91 g L–1 h–1 were obtained under phosphate limitation.45 Continuous fermentation has been optimized for the efficient production of mcl-PHA.32,33,35 It has been shown that the PHA content decreased with increased specific growth rate. Thus, a compromise between PHA content and productivity is required in a single-stage continuous process.43 A twostage continuous process was developed to overcome this limitation.46 Cell densities of 18 g L–1 with a PHA content of 63 w% and an overall volumetric productivity of 1.06 g L–1 h–1 were obtained.

16.2.2 Material properties Interestingly, the material properties of PHA are similar within one class as shown in Table 16.3. They are strongly dependent on the monomeric composition, in particular on the monomers’ side-chain and on the distance between the ester linkages in the backbone. In general, PHA is water insoluble, biodegradable, and biocompatible. It can be degraded at a moderate rate (3– 9 months) by many microorganisms into carbon dioxide and water using their own secreted PHA depolymerases.47 Its primary breakdown products, 3-hydroxyacids, are naturally found in animals and humans. Scl-PHAs are typically crystalline thermoplasts. The homopolymer PHB is a relatively stiff and brittle bioplastic, which can be processed by melt extrusion. PHA co-polymers composed of primarily HB with a fraction of longer chain monomers, such as HV, HHx or HO, are more flexible and less brittle thermoplasts. Their structures are shown in Fig. 16.5. They can be used in melt-extrusion processes to form a wide variety of products including containers, bottles, razors, and materials for food packaging. PHB and PHBV have been used as a water-proof film on the back of diaper sheets.48 PHBV is more flexible, more impact resistant and was marketed under the trade name Biopol™ by ICI/Zeneca and later on by Monsanto until 1995. Copolymers consisting of 3-hydroxybutyrate and a 3-hydroxyalkanoic acid with at least six carbon atoms, for instance PHBHx, were developed by Procter Table 16.3 Physical properties of short-chain-length (scl) and mediumchain-length (mcl) PHAs

Crystallinity [%] Glass transition temperature [°C] Melting point [°C] Elongation to break [%] Density [g/cm3] n.d.: non-detectable

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Mcl-PHA

40–80 –8–9 80–180 6–10 1.25

20–40 –60–14 n.d.–80 300–450 1.05

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O

O

O

O

O

O

O

O

n

m

Poly(3-hydroxybutyrate-co-3-hydroxyvalerate)

n

m

Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate)

O O

n

Poly(4-hydroxybutyrate)

16.5 Chemical structures of commercially important scl-PHAs: Poly(3hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), poly(3hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx) and poly(4hydroxybutyrate) (P4HB).

and Gamble and commercialized under the name Nodax. In analogy to PHBV, PHBHx is less crystalline than PHB and more flexible. Common mcl-PHAs are less crystalline and more elastomeric. Their properties largely depend on their side-chains. For instance, the glass transition temperature of a mcl-PHA containing linear chains up to seven carbon atoms and aromatic groups varies between –39°C and –6°C with an increasing content of aromatic side-chains.49 The physical properties of mcl-PHAs containing functional groups in the side-chains can be adapted through chemical modification. Thereby, the field of possible applications is considerably extended. Through the insertion of hydroxy groups in the side-chains of mclPHA, the polymer’s polarity and solubility in polar solvents could be drastically changed.50–52 Epoxidation and crosslinking of mcl-PHAs with unsaturated side-chains increased the polymer’s tensile strength and Young’s modulus by a factor of 4 and 39, respectively.53 Polyhedral oligomeric silsesquioxane (POSS) was linked to the side-chains of unsaturated mcl-PHAs thereby increasing the crystallinity and raising the melting point up to 120°C.54 Very little is known about lcl-PHA. It has been shown that the ability to polymerize long-chain-length PHAs is limited to Pseudomonas putida KTOY06.55

16.2.3 Recovery methods Recovery processes can be divided into two categories: solvent extraction and chemical digestion. In the former, PHA is extracted from biomass by using an organic solvent like methylene chloride. In the latter case the rest biomass is digested by applying chemicals like sodium hypochlorite or

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hydrolytic enzymes. Through cross-flow filtration or centrifugation, PHA is then separated from cell debris. The main problem of chemical digestion is severe degradation of the PHA, resulting in a reduction of the molecular weight. To certify polymers for medical applications, very demanding specifications have to be fulfilled. Considerable amounts of impurities like proteins, surfactants and endotoxins are not tolerated by regulatory agencies. The presence of bacterial endotoxins in medical polymers is one of the biggest concerns of suppliers.56 Up to now only few methods for depyrogenation of PHA have been were described in literature, although this aspect is very important for medical applications. More details are given in Section 16.4.

16.3

Chemical digestion of non-PHA biomass

Digestion of non-PHA biomass is performed in aqueous environments without or with only small amounts of organic solvents (Fig. 16.6a). The suspended cells are usually lysed to release the PHA granules, which form an aqueous suspension. For cell lysis, various techniques used in biotechnology can be applied as shown in Table 16.4. The granule envelope protects the PHA to a certain degree from chemical degradation. All the PHA-free biomass constituents are rendered water soluble by chemical modification. This can be done by using chemical agents like peroxides, hypochlorites and ozone or by using hydrolytic enzymes. After cell disruption and digestion of PHAfree biomass, PHA granules can be separated from cell debris by conventional methods like centrifugation or cross-flow filtration. While PHB granules have a density of about 1.2 g cm–3, mcl-PHA granules possess a density close to that of water57 and therefore do not settle in aqueous suspensions.58 Hence, ultracentrifugation or cross-flow filtration has to be applied.59,60 Although the PHA granules are not physically dissolved in H2O, they form a stable suspension (latex). Once the PHA granules are isolated, further purification steps usually follow. In the following we will focus the discussion on the first step, the degradation of residual biomass, because this is crucial for the further processing. Cell debris digestion with hypochlorite61–68 was one of the earliest methods for degrading PHA-free biomass. Most cell constituents are oxidized by hypochlorite and thus become water soluble. However, PHA was shown to become severely degraded with a decrease in molecular weight Mw from e.g. 1200 kDa to 600 kDa under optimized conditions.61–63,69 In addition, it is difficult to eliminate residues of sodium hypochlorite completely in the subsequent steps. Hypochlorite digestion could be improved by combining it with surfactants or organic solvents and resulted in PHA of higher purity with less polymer degradation.62,69,70 Alternative oxidizing agents like hydrogen peroxide or ozone were investigated to oxidize PHA-free biomass. Peroxides were especially useful

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PHA + cell debris

Chemical digestion

PHA + dig. cell debris

Ultrafiltration or centrifugation

PHA latex Wash & drying

PHA (a)

Cell broth

Centrifugation or filtration

Wet biomass

Drying

Dry biomass

Solvent extraction

PHA solution + cell debris

Filtration or centrifugation PHA

Wash & drying

Wet PHA

Precipitation or evaporation

(b)

16.6 PHA recoveries through chemical digestion (a) and solvent extraction (b).

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PHA solution

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Cell broth

Cell disruption

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Table 16.4 Methods to lyse bacterial cells for the recovery of PHA Physical methods Bead milling Homogenization Ultrasonication Thermal shock Supercritical fluid treatment

Chemical methods [97, 115] [95, 97] [116] [117] [118, 119]

Reducing or oxidizing agents Detergents Lytic enzymes Solvents (osmotic shock) pH shock

[97, 109, 89] [108] [111, 120] [115]

to degrade nucleic acids and to decrease the viscosity of the lysate.71–73 In combination with enzymes and surfactants, the efficiency of rest biomass degradation with peroxides could be increased.74 Treatment of PHA containing biomass or partially purified PHA with ozone yielded an enhanced level of purity in combination with other purification steps. The resulting polymer was basically odor-free.75 Furthermore, coloration of the PHA was drastically reduced with ozone or peroxide. It has not clearly been described in literature if PHA was degraded by hydrogen peroxide or ozone, but a certain degradation has to be expected due to the strong reactivity of these agents. Simple alkaline digestion of suspended biomass provided PHB with a purity of 85% or higher, but a minor degradation of PHA has also been observed.76 Combination with chelates and surfactants improved cell disruption and solubilisation of the rest biomass, and therefore PHA of higher purity was obtained. The use of enzymes is an alternative to harsh chemical agents. Cell fragments can be degraded more selectively and the degradation of PHA is negligible. Enzyme cocktails consisting of various hydrolytic enzymes like proteinases, nucleases, phospholipases, lysozymes and others in combination with surfactants and chelate formers have been applied.16,74,77–80 Combined with heat treatment, PHA purities of up to 95% could be achieved.81 Heat accelerated the degradation of residual biomass, while surfactants and chelates improved the solubility of all components. Mcl-PHA with a purity of 93 w% was obtained by enzymatic digestion combined with ultrafiltration.60 Generally the isolation of the PHA-granules is much easier with bacteria that have a high PHA content (>60% in dry mass). Only one or two steps are necessary to achieve a purity of around 90%. If the PHA content is below 60%, the separation process is more complicated. For this case a combined method with enzymes and reducing agents like sodium dithionite was effective.82 Cell components like proteins, nucleic acids and polysaccharides were decomposed without drastically changing PHA properties. To avoid the problem of a dramatic increase in viscosity caused by the liberation of DNA, a nuclease-encoding gene was integrated into the genome of PHA producing bacteria.83 Thereby the amount of enzymes or chemicals necessary for the digestion of the rest biomass was reduced. The lysate

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viscosity was significantly reduced without affecting the PHA production or the strain stability. Major disadvantages of the enzymatic digestion are the high costs of hydrolytic enzymes and the additional purification steps necessary to reach a high degree of purity. Moreover surfactants can hardly be separated from PHA. The purification of PHA granules would be much more efficient if they were secreted into the medium. Separation from the cells could be done by centrifugation without digesting the PHA-free biomass. The production of extracellular PHA has been proposed.84

16.3.1 Standard method: Solvent extraction and precipitation of PHA in non-solvents As discussed above, the alternative method to chemical or enzymatic digestion of the PHA-free biomass is the selective extraction of PHA from biomass with an organic solvent. Most recovery procedures based on the extraction with organic solvents follow the steps shown in Fig. 16.6b. Bacterial cells are collected by centrifugation or filtration and dried to remove water that inhibits an effective extraction. Pre-treating dry biomass with methanol can be effective to remove some of the lipids and coloring impurities.85,86 PHA is extracted from the dried biomass with an organic solvent under stirring and in some cases heating. The resulting suspension is filtered or centrifuged to remove particulate cell debris and subsequently PHA is either precipitated with a non-solvent, or obtained by evaporation of the solvent. In some cases the crude PHA thus obtained is washed with a non-solvent. Repeated extraction, precipitation and washing steps result in a higher purity. Solvent extractions normally use large amounts of solvents, typically 5– 20 times the dry weight of biomass and similar amounts of non-solvents for precipitation. Solvent recycling is energy consuming, especially for solvents with high boiling points. The high viscosity of even diluted PHA solutions (e.g. 5% w/v) impedes extensive solvent savings and limits economical optimization. Soxhlet extraction was used to work with reduced volumes of solvent.87 The solubility of PHA is strongly dependent on the polymer composition, the molecular weight and on the temperature and pressure.88 The extraction of scl-PHA was usually performed with chlorinated solvents like methylene chloride, chloroform, 1,2-dichloroethane, 1,1,2-trichloroethane and 1,1,2,2-tetrachloroethane89–97 because most chlorinated solvents are capable of dissolving PHB at a relatively high concentration, but only a little of the rest biomass is dissolved as well. Preferentially the extraction was carried out at elevated temperatures to increase the solubility and to reduce the viscosity of the polymer solution and the extraction time. To reach high temperatures without evaporation of the solvent, pressurization has been applied.98 PHA degradation at high extraction temperatures has been observed, especially when water was present in solution. At temperatures above 200°C

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degradation occurs also in the absence of water by a non-radical random chain scission reaction (cis-elimination).99 Recently, it has been shown that even at moderate temperatures thermal degradation occurs via E1cB mechanism if carboxylate groups are present 100. To precipitate the dissolved polymer, various non-solvents like methanol, ethanol, water, or ether have been used.20,91,94,101,102 Thereby, ether could only be applied for scl-PHA as it dissolves mcl-PHAs. PHB with a purity of 99% could be obtained with precipitation in water.103 Mixtures of chlorinated solvents with a non-solvent were applied to extract PHB from biomass at high temperatures and to precipitate the polymer by cooling to room temperature.93 Azeotrope-building chlorinated solvents (e.g. 1,1,2-trichloroethane) were used for simultaneous azeotropic distillation and PHA extraction from an aqueous suspension of microorganisms.96 Thereby water was removed from the cell suspension as a minimum boiling azeotrope. The remaining chlorinated solvent was used to extract PHA. Hence a preceding drying of the cell suspension could be avoided. In an effort to circumvent the disadvantages of halogenated solvents, which are known to pose health risks and environmental problems, alternative solvents like cyclic carbonic acid esters,104 methyl lactate,105 ethyl lactate,105,106 acetic acid,107–109 acetic acid anhydride,110,111 n-methylpyrrolidone,112 tetrahydrofuran 113 and mixtures of non-halogenated solvents 114 were investigated for the extraction of PHB. Co-polymers with PHB were also extracted by non-chlorinated solvents like acetone,85 ethyl acetate,115 butyl acetate,116 methyl isobutylketone116 and cyclo-hexanone.116 Mcl-PHA has the advantage of being soluble in a broader spectrum of solvents than PHB. Even at room temperature, typical mcl-PHAs are soluble in acetone, THF or diethyl ether. A patent from Firmenich SA describes an extraction method with non-halogenated solvents at room temperature tailored for polyhydroxyoctanoate.117 More recently, solvents like n-hexane, 2-propanol and acetone were used for extracting mcl-PHA from biomass.86,118 A combined extraction-filtration method has been described using a circular filtration system,119 where the aqueous slurry was diafiltrated and an organic solvent like acetone was continuously added. In the beginning the bacterial cells were rejected, but at a certain acetone to water ratio of about 9:1 (w/w) the cells were lysed and mcl-PHA was dissolved and passed the filtration membrane.

16.3.2 Temperature controlled extraction and precipitation The recycling of solvent and non-solvent mixtures is time and energy consuming. The utilization of non-solvents could be avoided by temperature induced extraction and precipitation.98,118 The solvent system has to be properly

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adapted to the PHA of interest because the solubility of different PHAs varies. PHA is extracted at elevated temperature and precipitated by cooling the polymer solution to a lower temperature. The solvent is separated from the gel-like polymer by decantation. This procedure was effective by using n-hexane as solvent for the extraction of PHO. Optimal conditions for PHO were an extraction temperature of 50°C and a precipitation temperature of 5°C resulting in a purity of more than 95%.118 The endotoxic contamination was 15 EU g–1 PHA. The analogous approach with 2-propanol resulted in a lower purity and recovery for PHO, because more polar impurities were coextracted and co-precipitated. Temperature controlled precipitation can alter the molecular weight distribution because high molecular weight polymers are less soluble than those of low molecular weight. Consequently, this separation effect could be used when low molecular weight PHAs have to be excluded, concomitantly reducing the molecular weight polydispersity.

16.3.3 Special approaches for the recovery of PHA Supercritical fluid extraction (SFE) with carbon dioxide has been proposed for the recovery of PHA.56 After extraction, carbon dioxide evaporates immediately and a drying step is not required. However, published data are conflicting. According to Khosravi-Darani,120 pure PHB is soluble up to 8.01 g L–1 in supercritical CO2 at a temperature of 348 K and a pressure of 355 bar. In contrast to these results, Seidel and Hampson showed that lipids, pigments and ubiquinones were extracted from biomass containing PHB or PHO without dissolving the polymer under similar conditions. In these reports PHA was extracted from the purified biomass with an organic solvent, e.g. chloroform.121,122 Consequently, it remains uncertain whether SFE can be efficiently used to extract PHA from biomass. A process for recovering PHA from biological material by comminution and air classification has been patented.123,124 In principle no chemicals or solvents are needed for this process. The dried biomass was defatted and ground so that the diameter of most particles was below 100 µm. An air stream was used to suspend the particles and classify them according to their weight or size. At appropriate stream velocities the particles were separated into a coarse and a fine fraction. The fine fraction was then subjected to further purification steps to reach a PHA purity of 80% or higher. So far, it has not been described whether this process is successfully applied for the recovery of PHA on a larger scale. Furthermore dissolved-air flotation was recently used to separate mclPHA granules from the fermentation broth.125 Flotation is potentially cheap and is common in wastewater treatment. Flotation separates particles according to their affinity to the air/liquid interface which is generated by bubbling air (or another gas) through a liquid phase (e.g. water). Hydrophobic particles

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are transported with the air-bubbles to the surface, whereas more hydrophilic ones remain in the aqueous phase. The cells were pre-treated with enzymes to release the PHA granules. Selective aggregation and flotation of mcl-PHA granules could be triggered by adjusting the pH at around 3.5. A purity of 86% was obtained after three consecutive batch flotation steps.

16.4

Purification of PHA

For applications in the medical field, PHA of high purity is needed. In particular, biologically active contaminants like proteins and lipopolysaccharides (LPS) have to be reduced to a very low level as they could induce immunoreactions. The US Food and Drug Administration (FDA) requires the endotoxin content of medical devices not to exceed 20 US Pharmacopeia (USP) endotoxin units (EU) per device, except for those devices that are in contact with the cerebrospinal fluid. In this case the content must not exceed 2.15 USP EU per device. LPS act as endotoxins whereby minute quantities can have severe effects in contact with blood and trigger immunoreactions. Particularly for PHA derived from fermentation of Gramnegative bacteria, contamination with endotoxins is a serious problem, as LPS are part of the outer membrane. During cell lysis and product recovery, LPS are liberated from the outer membrane and contaminate PHA. Therefore, PHA for use in medical devices has to be carefully purified from such endotoxins. Standard techniques used for the purification of PHA are re-dissolution and precipitation, washing with a non-solvent, purification by chromatography, treatments with chemical agents and filtration. The purity of PHA can also be increased by washing the biomass before extracting PHA or by aqueous digestion to remove major impurities as described before.

16.4.1 Sources and characterization of contamination The spectrum of possible PHA-contaminants from PHA-free biomass is broad. But basically, the extraction solvent determines which impurities will be carried over. For instance, proteins and DNA have been frequently detected when PHA was recovered by aqueous chemical digestion. Solvent extraction with non-polar solvents or with solvents of average polarity is more susceptible to co-extract lipids and coloring substances. Notably polar organic solvents like acetone and 2-propanol co-extract chromophores and give the polymer a yellow to brownish color. Lipopolysaccharides have been detected with aqueous digestion as well as with solvent extraction. They are soluble in water, but their amphiphilic character renders them to some extent soluble in non-polar solvents. Besides lipids, proteins and lipopolysaccharides, also antifoam agents

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Natural-based polymers for biomedical applications 0.8

Absorbance

0.6 1 0.4 2

0.2

3 0.0

4 200

300

400 Wavelength (nm)

500

600

16.7 UV spectra of crude extracted and purified mcl-PHA in chloroform. 1: acetone extract precipitated in methanol, 2-4: additional cycles of dissolution and precipitation (1-3 times).122

from fermentations, surfactants and hydrolytic enzymes from the purification procedure are contaminants of the polymer. The most common impurities found in PHA are summarized in Table 16.5. In summary, the nature of contaminants is determined by the biosynthesis as well as by the downstream processing.

16.4.2 Methods to purify extracted PHA Repeated dissolution and precipitation was commonly applied to reach a purity of close to 100%. The efficiency of this method is dependent on several parameters like the concentration of the polymer solution and the temperature, but usually, large amounts of non-solvents were used. Jiang et al.86 showed that the concentration of UV absorbing molecules is considerably decreased by dissolution of mcl-PHA in acetone and precipitation in cold methanol, as shown in Fig. 16.7. Scl-PHA was repeatedly dissolved in chloroform and precipitated in ethanol to remove biologically active substances. Traces of fatty acids from C6 to C18 were eliminated and the hemocompatibility increased.126 The authors propose that these fatty acids originated from lipopolysaccharides. Temperature controlled dissolution and precipitation in 2-propanol was used to purify certain mcl-PHA like PHO. A purity of nearly 100% and an endotoxicity of below 10 EU/g polymer was obtained with this method.118 To reduce the endotoxin content further to acceptable values, e.g. < 10 EU/ gram of PHA, a treatment with an oxidizing agent such as hydrogen peroxide or benzoyl peroxide was used successfully.127 Destruction of LPS by applying basic conditions was also successful.128,129 The concentration of base and the treatment time (Fig. 16.8) are crucial for an ideal detoxification. Otherwise,

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Table 16.5 Common contaminants found in PHA Amount in PHA

Ref.

Recovery method

Microorganism

PHA

Lipids

0.1-3 mol%a

[161]

Ethanol-KOH wash + chloroformethanol extraction

R. eutropha B5786

PHB, PHBV

Proteins

0-9 w% N.d N.d

[95, 162] [156] [86]

Homogenization + centrifugation Extraction Aq. digestion

E. coli M. rhodasianum N.d.

PHB PHB PHO

DNA

0.03-2.2 w% 0.6-2.2 w%

[95] [162]

Homogenization + centrifugation Homogenization + centrifugation

E. coli E. coli

PHB PHB

LPS

>120 EU/g 1-104 EU/g

[86] [163]

N.d. NaOH-digestion

N.d. E. coli

Commercial PHB PHB

Antifoam agents

< 1 w%

Unpublished results

Chloroform-ethanol extraction

P. putida GPo1

PHO

SDS

N.d.

[111]

SDS solubilisation + enzymatic treatment

P. putida GPo1

PHO

UV absorbers

N.d.

[122]

Acetone extraction

P. putida KT2440

mcl-PHA

Notes: N.d. not defined a Relative to the PHA monomers

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Contaminant

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Endotoxin level (EU/g PHB)

106 105 104 103 102

10

1 0

1

2 3 4 5 NaOH digestion time (h)

10

16.8 Endotoxic activity of PHB recovered by 1.2N NaOH digestion at 30°C for various durations.163

the destruction of LPS is incomplete or PHA is depolymerized and its molecular weight reduced as previously mentioned. Further methods can be used for purification such as washing of PHA with a non-solvent, purification by chromatography, filtration and treatment with endotoxin removing agents; however, accurate data about their efficacy are not available. For the protein purification in aqueous solutions, e.g. cationic endotoxin removal agents have been shown to be very successful.130

16.5

Potential applications of PHA in medicine and pharmacy

PHA has the potential to become an important compound for medical applications.56,131 Biocompatibility and slow biodegradability are thereby essential properties. The changing PHA composition also allows favorable mechanical properties as shown in Table 16.3. In vitro cell experiments and in vivo studies have focused on PHB, PHBV, P4HB, PHBHx and PHO. An overview of the potential applications of PHA is given in Table 16.6. In the following we discuss the applications in the field of drug delivery and tissue engineering. The discussion is restricted to these fields, because the most promising research was done in these areas. © 2008, Woodhead Publishing Limited

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Table 16.6 Potential applications of PHA in medicine and pharmacy Type of application

Products

Type of PHA

Wound management

Sutures, skin substitutes, nerve cuffs, surgical meshes, staples, swabs

Scl, mcl

Vascular system

Heart valves, cardiovascular fabrics, pericardial patches, vascular grafts

Mcl

Urology

Urological stents

Scl, mcl

Orthopaedy

Scaffolds for cartilage engineering, spinal cages, bone graft substitutes, meniscus regeneration, internal fixation devices (e.g. screws)

Scl, (mcl)

Dental

Barrier material for guided tissue regeneration in periodontosis

Scl, mcl

Computer assisted tomography and ultrasound imaging

Micro- and nanospheres for anticancer therapy

Scl, mcl

Drug delivery

Chemoembolizing agents, microand nanospheres for anticancer therapy

Scl, mcl

16.5.1 PHA as drug carrier PHAs became candidate material as drug carriers in the early 1990s due to their inherent biocompatibility.132 Microspheres of PHB loaded with rifampicin were investigated for their use as a chemoembolizing agent (agent for the selective occlusion of blood vessels).133,134 The drug release of all microspheres was very rapid, with almost 90% of the drug released within 24 h. The drug release rate could be controlled by the drug loading and the particle size. A similar behavior was described by Sendil and coworkers for PHBV supplemented with tetracycline.135 PHBV was further investigated as an antibiotic-loaded carrier to treat implant-related and chronic osteomyelitis.136 The antibiotic sulbactamcefoperazone was integrated into PHBV rods and implanted into a rabbit tibia that was artificially infected by S. aureus. The infection subsided after 15 days and was nearly completely healed after 30 days. In search of an efficient transdermal drug delivery system, a PHO-based system with a polyamidoamine dendrimer was examined. Tamsulosin was used as the model drug. The dendrimer was found to act as a weak permeability enhancer. By adding the dendrimer, the dendrimer-containing PHA matrix achieved the clinically required amount of tamsulosin permeating through the skin model.137

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16.5.2 PHA as scaffold material in tissue engineering PHBV was chosen as a temporary substrate for growing retinal pigment epithelium cells as an organized monolayer before their subretinal transplantation. The surface of the PHBV film was hydrophilized by oxygen plasma treatment to increase the attachment of D407 cells to the polymer surface. The cells grew to confluency as an organized monolayer. Hence, PHBV films can be used as temporary substrates for subretinal transplantation to replace diseased or damaged retinal pigment epithelium.138 An interesting approach is the implantation of biodegradable supporting scaffolds that are seeded with tissue-engineered cells. This approach was exemplified by Sodian and co-workers139 who used PHO and P4HB for the fabrication of a tri-leaflet heart valve scaffold. A porous surface was achieved with the salt leaching technique, resulting in pore sizes between 80 and 200 µm. The scaffold was seeded with vascular cells from ovine carotid artery and subsequently tested in a pulsatile flow bioreactor. The cells formed a confluent layer on the leaflets. In another study, the native pulmonary leaflets were resected with the use of cardiopulmonary bypass, and segments of pulmonary artery were replaced by autologous cell-seeded heart valve constructs. All animals survived the procedure without receiving any anticoagulation therapy. The tissue engineered constructs were covered with tissue and no thrombus formation was observed. It was concluded that tissue engineered heart valve scaffolds fabricated from PHO can be used for implantation in the pulmonary position with an appropriate function for 120 days in lambs.6 The same group demonstrated that PHO and P4HB have thermoprocessible advantages over PGA, which has better properties for ovine vascular cell growth.139,140 Vascular smooth muscle cells and endothelial cells from ovine carotid arteries were seeded on P4HB scaffolds to study autologous tissue engineered blood vessels in the descending aorta of juvenile sheep. Up to three months after implantation, grafts were fully patent, without any signs of dilatation, occlusion or intimal thickening. A confluent luminal endothelial cell layer was observed. In contrast, after six months, the graft displayed significant dilatation and partial thrombus formation, most likely caused by an insufficient elastic fiber synthesis.141 PHBHx was found to be a suitable biomaterial for osteoblast attachment, proliferation and differentiation from bone marrow cells. The cells on PHBHx scaffolds presented typical osteoblast phenotypes: round cell shape, high alkaline phosphatase (ALP) activity, strong calcium deposition, and fibrillar collagen synthesis. After incubation for ten days, cells grown on PHBHx scaffolds were approximately 40% more than those on PHB scaffolds and 60% more than those on PLA scaffolds. ALP activity of the cells grown on PHBHx scaffolds was up to about 65 U g–1 scaffold, 50% higher than that of

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PHB and PLA, respectively.142 Similarly, it was observed that chondrocytes isolated from rabbit articular cartilage proliferated better on PHB scaffolds blended with PHBHx than on pure PHB scaffolds. Chondrocytes proliferated on the PHB-PHBHx scaffold and preserved their phenotype up to 28 days.143

16.6

Conclusions and future trends

Polyhydroxyalkanoates with a wide range of physical properties are accessible through biosynthesis in bacteria. It has been shown that they have a potential in several medical applications. Unfortunately, inappropriate downstream processing of the polymer resulted in contamination of PHAs by bacterial cell compounds and therefore may have affected first studies in a negative way. Improved purification methods have been described in recent years which were successful in reducing pyrogenic contaminations. Recently, a type of PHA, P4HB, obtained the approval of the US Food and Drug Administration for application as a suture material. It is to be expected that more PHAs will follow because material properties of PHAs can already be tailored for particular applications during biosynthesis, or later on by chemical and physical modifications.

16.7

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78 Kim K, Sonn Y, Lee M and Jung S, Preparation process of poly-3-hydroxybutyrate, Patent KR9609065, 1996. 79 Ramsay B A, Ramsay J, Berger E, Chavarie C and Braunegg G, Separation of poly(β-hydroxyalkanoic acid) from microbial biomass, Patent US5110980, 1992. 80 Yamamoto O, Myata Y and Yanagi S, Separation and purification of poly(hydroxyalkanoates) from microorganisms using surfactants, Patent JP07079787, 1995. 81 deKoning G J M and Witholt B, A process for the recovery of poly(hydroxyalkanoates) from Pseudomonads .1. Solubilization, Bioproc Eng, 1997, 17(1), 7–13. 82 Schumann D and Müller R A, Method for obtaining polyhydroxyalkanoates (PHA) and the copolymers thereof, Patent WO0168892, 2003. 83 Boynton Z L, Koon J J, Brennan E M, Clouart J D, Horowitz D M, Gerngross T U and Huisman G W, Reduction of cell lysate viscosity during processing of poly(3-hydroxyalkanoates) by chromosomal integration of the staphylococcal nuclease gene in Pseudomonas putida, Appl Environ Microb, 1999, 65(4), 1524– 1529. 84 Sabirova J S, Ferrer M, Lunsdorf H, Wray V, Kalscheuer R, Steinbüchel A, Timmis K N and Golyshin P N, Mutation in a ‘tesB-like’ hydroxyacyl-coenzyme A-specific thioesterase gene causes hyperproduction of extracellular polyhydroxyalkanoates by Alcanivorax borkumensis SK2, J Bacteriol, 2006, 188(24), 8452–8459. 85 Gorenflo V, Schmack G, Vogel R and Steinbüchel A, Development of a process for the biotechnological large-scale production of 4-hydroxyvalerate-containing polyesters and characterization of their physical and mechanical properties, Biomacromolecules, 2001, 2(1), 45–57. 86 Jiang X, Ramsay J A and Ramsay B A, Acetone extraction of mcl-PHA from Pseudomonas putida KT2440, J Microbiol Meth, 2006, 67(2), 212–219. 87 Valappil S P, Peiris D, Langley G J, Hemiman J M, Boccaccini A R, Bucke C and Roy I, Polyhydroxyalkanoate (PHA) biosynthesis from structurally unrelated carbon sources by a newly characterized Bacillus spp, J Biotechnol, 2007, 127(3), 475– 487. 88 Terada M and Marchessault R H, Determination of solubility parameters for poly(3hydroxyalkanoates), Int J Biol Macromol, 1999, 25(1–3), 207–215. 89 Baptist J N, Process for preparing PHB, Patent US 3044942, 1962. 90 Holmes P A, Wright L F, Alderson B and Senior P J, Extraction of poly(3hydroxybutyric acid) from microbial cells, Patent EP15123, 1980. 91 Numazawa R, Miyamori T, Sakimae A and Onishi H, Separation and purification of poly(β-hydroxybutyric acid) from cell extracts, Patent JP62205787, 1987. 92 Ramsay J A, Berger E, Voyer R, Chavarie C and Ramsay B A, Extraction of poly3-hydroxybutyrate using chlorinated solvents, Biotechnol Techn, 1994, 8(8), 589– 594. 93 Schmidt J, Schmiechen H, Rehm H and Trennert M, Verfahren zur Gewinnung von PHB aus getrockneten Biomassen, Patent DD239609, 1986. 94 Stageman J F, Extraction process, Patent US4562245, 1985. 95 Vanlautem N and Gilain J, Process for separating poly(β-hydroxybutyrates) from a biomass, Patent US4310684, 1982. 96 Vanlautem N and Gilain J, Process for extracting poly(β-hydroxybutyrates) by means of a solvent from an aqueous suspension of microorganisms, Patent US4705604, 1987. 97 Barham P J, Extraction of poly(β-hydroxybutyric acid), Patent EP58480, 1982.

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98 Kurdikar D L, Strauser F E, Solodar A J and Paster M D, High temperature PHA extraction using PHA-poor solvents, Patent WO9846783, 1998. 99 Lee M Y, Lee T S and Park W H, Effect of side chains on the thermal degradation of poly(3-hydroxyalkanoates), Macromol Chem Phys, 2001, 202(7), 1257– 1261. 100 Kawalec M, Adamus G, Kurcok P, Kowalczuk K, Foltran I, Focarete M L and Scandola M, Carboxylate induced degradation of PHB, Biomacromolecules, 2007, 8(4), 1053–1058. 101 Baptist J N, Process for preparing PHB, Patent US3036959, 1962. 102 Holmes P A, A process for the extraction of PHB from microbial cells, Patent EP0015123, 1980. 103 Hrabak O, Industrial production of poly(β-hydroxybutyrate), FEMS Microbiol Rev, 1992, 103(2–4), 251–255. 104 Agroferm, Use of cyclic carbonic acid esters as solvent for poly(β-hydroxybutyric acid), Patent US4140741, 1977. 105 Metzner K, Sela M and Schaffer J, Agents for extracting polyhydroxyalkanoic acids, Patent WO9708931, 1997. 106 Sela M and Metzner K, Use of ethyl lactate as an extractant for poly(hydroxyalkanoic acids), Patent DE19533459, 1996. 107 Rapthel I, Lehmann O, Runkel D, Mayer T and Rauchstein K D, Manufacture of colorless polyhydroxyalkanoates by extraction of bacterial biomass with acetic acid containing β-butyrolactone, Patent DE4215860, 1993. 108 Rapthel I, Lehmann O, Runkel D, Mayer T, Rauchstein K D and Schaffer J, Manufacture of colorless polyhydroxyalkanoates by extraction of bacterial biomass with acetic acid containing acetic anhydride, Patent DE4215861, 1993. 109 Runkel D, Lehmann O, Mayer T, Rauchstein K D and Schaffer J, Manufacture of colorless polyhydroxyalkanoates by extraction of bacterial biomass with acetic acid, Patent DE4215862, 1993. 110 Runkel D, Lehmann O, Mayer T, Rauchstein K D and Schaffer J, Manufacture of colorless polyhydroxyalkanoates by extraction of bacterial biomass with acetic anhydride, Patent DE4215864, 1993. 111 Schmidt J, Biedermann W and Schmiechen H, Extraction of poly(β-hydroxybutyric acid) from bacterial biomass, Patent DD229428, 1985. 112 Schumann D and Mueller R A, Method for obtaining polyhydroxyalkanoates (PHA) or the copolymers thereof, Patent WO2001068892, 2001. 113 Matsushita H, Yoshida S and Tawara T, Extraction of poly(3-hydroxybutyric acid) from microorganisms, Patent JP07079788, 1995. 114 Noda I and Schechtman L A, Solvent extraction of polyhydroxyalkanoates from biomass, Patent WO9707230, 1997. 115 Chen G Q, Zhang G, Park S J and Lee S Y, Industrial scale production of poly(3hydroxybutyrate-co-3-hydroxyhexanoate), Appl Microbiol Biot, 2001, 57(1-2), 50– 55. 116 Walsem H, Zhong L and Shih S, Polymer extraction methods, Patent US2004/ 013204, 2004. 117 Ohleyer E, Extraction of poly(β-hydroxyoctanoate) from microbial biomass, Patent WO9311656, 1993. 118 Furrer P, Panke S and Zinn M, Efficient recovery of low endotoxin medium-chainlength poly([R]-3-hydroxyalkanoate) from bacterial biomass, J Microb Meth, 2007, 69(1), 206–213.

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119 Horowitz D, Methods for purifying polyhydroxyalkanoates, Patent WO2000068409, 2000. 120 Khosravi-Darani K, Vasheghani-Farahani E, Yamini Y and Bahramifar N, Solubility of poly(β-hydroxybutyrate) in supercritical carbon dioxide, J Chem Eng Data, 2003, 48(4), 860–863. 121 Seidel H, Voigt B, Roethe K P, Rosahl B and Mothes S, Supercritical fluid extraction in polyhydroxybutyrate and ubiquinone extraction for bacteria, Patent DD294280, 1991. 122 Hampson J W and Ashby D, Extraction of lipid-grown bacterial cells by supercritical fluid and organic solvent to obtain pure medium chain-length polyhydroxyalkanoates, J Am Oil Chem Soc, 1999, 76(11), 1371–1374. 123 Noda I, Process for recovering polyhydroxyalkanoates using air classification, Patent US5849854, 1998. 124 Noda I, Process for recovering polyhydroxyalkanoates using air classification, Patent WO9533064, 1995. 125 van Hee P, Elumbaring A, van der Lans R and van der Wielen L A M, Selective recovery of polyhydroxyalkanoate inclusion bodies from fermentation broth by dissolved-air flotation, J Colloid Interf Sci, 2006, 297(2), 595–606. 126 Sevastianov V I, Perova N V, Shishatskaya E I, Kalacheva G S and Volova T G, Production of purified polyhydroxyalkanoates (PHAs) for applications in contact with blood, J Biomater Sci Polym Ed, 2003, 14(10), 1029–1042. 127 Williams S F, Martin D P, Gerngross T and Horowitz D M, Removing endotoxin with an oxdizing agent from polyhydroxyalkanoates produced by fermentation, Patent US6245537, 2001. 128 Sevastianov V I, Perova N V, Sihshatskaya E I, Kalacheva G S and Volova T G, Production of purified polyhydoxyalkanoate (PHAs) for applications in contact with blood, J Biomater Sci Polymer Edn, 2003, 14(10), 1029–1042. 129 Lee S Y, Choi J I, Han K and Song J Y, Removal of endotoxin during purification of poly(3–hydroxybutyrate) from Gram-negative bacteria, Appl Env Microb, 1999, 65(6), 2762–2764. 130 Zhang J P, Qun Wang T R S, William E, Hurst and Sulpizio T, Endotoxin removal using a synthetic adsorbent of crystalline calcium silicate hydrate, Biotechnol Progr, 2005, 21, 1220–1225. 131 Zinn M, Witholt B and Egli T, Occurrence, synthesis and medical application of bacterial polyhydroxyalkanoate, Adv Drug Del Rev, 2001, 53(1), 5–21. 132 Pouton C W and Akhtar S, Biosynthetic polyhydroxyalkanoates and their potential in drug delivery, Adv Drug Deliver Rev, 1996, 18(2), 133–162. 133 Kassab A C, Piskin E, Bilgic S, Denkbas E B and Xu K, Embolization with polyhydroxybutyrate (PHB) microspheres: In-vivo studies, J Bioact Compat Polym, 1999, 14(4), 291–303. 134 Kassab A C, Xu K, Denkbas E B, Dou Y, Zhao S and Piskin E, Rifampicin carrying polyhydroxybutyrate microspheres as a potential chemoembolization agent, J Biomater Sci Polym Ed, 1997, 8(12), 947–961. 135 Sendil D, Gursel I, Wise D L and Hasirci V, Antibiotic release from biodegradable PHBV microparticles, J Control Rel, 1999, 59(2), 207–217. 136 Yagmurlu M F, Korkusuz F, Gursel I, Korkusuz P, Ors U and Hasirci V, Sulbactamcefoperazone polyhydroxybutyrate-co-hydroxyvalerate (PHBV) local antibiotic delivery system: In vivo effectiveness and biocompatibility in the treatment of implant-related experimental osteomyelitis, J Biomed Mater Res, 1999, 46(4), 494–503. © 2008, Woodhead Publishing Limited

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137 Wang Z X, Itoh Y, Hosaka Y, Kobayashi I, Nakano Y, Maeda I, Umeda F, Yamakawa J, Kawase M and Yagi K, Novel transdermal drug delivery system with polyhydroxyalkanoate and starburst polyamidoamine dendrimer, J Biosci Bioeng, 2003, 95(5), 541–543. 138 Tezcaner A, Bugra K and Hasirci V, Retinal pigment epithelium cell culture on surface modified poly(hydroxybutyrate-co-hydroxyvalerate) thin films, Biomaterials, 2003, 24(25), 4573–4583. 139 Sodian R, Sperling J S, Martin D P, Egozy A, Stock U, Mayer J E and Vacanti J P, Fabrication of a trileaflet heart valve scaffold from a polyhydroxyalkanoate biopolyester for use in tissue engineering, Tissue Eng, 2000, 6(2), 183–188. 140 Sodian R, Hoerstrup S P, Sperling J S, Martin D P, Daebritz S, Mayer J E and Vacanti J P, Evaluation of biodegradable, three-dimensional matrices for tissue engineering of heart valves, Asaio J, 2000, 46(1), 107–110. 141 Opitz F, Schenke-Layland K, Cohnert T U, Starcher B, Halbhuber K J, Martin D P and Stock U A, Tissue engineering of aortic tissue: dire consequence of suboptimal elastic fiber synthesis in vivo, Cardiovas Res, 2004, 63(4), 719–730. 142 Wang Y W, Wu Q O and Chen G Q A, Attachment, proliferation and differentiation of osteoblasts on random biopolyester poly(3-hydroxybutyrate-co-3hydroxyhexanoate) scaffolds, Biomaterials, 2004, 25(4), 669–675. 143 Deng Y, Lin X S, Zheng Z, Deng J G, Chen J C, Ma H and Chen G Q, Poly(hydroxybutyrate-co-hydroxyhexanoate) promoted production of extracellular matrix of articular cartilage chondrocytes in vitro, Biomaterials, 2003, 24(23), 4273–4281.

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17 Electrospinning of natural proteins for tissue engineering scaffolding P. I. L E L K E S, M. L I, A. P E R E T S, L. L I N, J. H A N and D. W O E R D E M A N, Drexel University, USA

17.1

Introduction

Tissue Engineering (TE), according to a consensus definition, is ‘an emerging multidisciplinary field involving biology, medicine, and engineering that aims at restoring, maintaining, or enhancing tissue and organ function’ (MATES, 2007). In the past, much of TE revolved around the biomaterials-centered approaches toward engineering ‘biocompatible scaffolds’. For the most part porous scaffolds were made of degradable synthetic biomaterials such as polyglycolides, on/in which cells were seeded and cultured in vitro. The expectation was that this combination of biodegradable scaffolds and cells would yield functional tissue constructs, which, upon implantation by surgeons, would eventually fulfill the aim of replacing or augmenting the functions of the damaged or diseased tissues in situ. Unfortunately, this rather generic ‘engineer-centric’ approach has largely failed to live up to the tremendous potential and promise (Nerem, 2006). Why? One potentially provocative answer to this question may be that the early phase of ‘tissue engineering’ was guided by too much emphasis on ‘engineering’ and too little on ‘tissue’, and by the expectation of instant pecuniary gratification and entrepreneurial success. More recent stratagems recognize the importance of approaching TE not only from the engineering aspect, but rather to focus more on the interdisciplinary aspects related to cells/tissues (Hunziker et al., 2006; Mikos et al., 2006). Specifically, the strategic plan of the Multi-Agency Tissue Engineering Science Interagency Working Group (MATES-IWG) formulated as two of its key strategic priorities to ‘obtain a molecular-level understanding of the basic physical, chemical, and molecular biological conditions that direct cells to assemble into and maintain cellular communities and functional 3-D tissues’ and to ‘develop design principles for new materials based on a physical and quantitative understanding of how cells respond to molecular signals and integrate multiple inputs to generate a given response in their physiological environment’ and ‘test new matrices for biocompatibility and successful integration into relevant hosts or in vitro’ (MATES, 2007). 446 © 2008, Woodhead Publishing Limited

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Indeed, recent successful approaches toward TE encompass a strong tissue/ biological component, focusing on cell-, molecular- and developmental biological aspects of the multitude of cells of the body; specifically stem or progenitor cells, and the plethora of specific (signaling) biomolecules that contribute to the formation and function of any given tissue and organ (Stenn and Cotsarelis, 2005; Ingber et al., 2006; Polak and Bishop, 2006; Carlson et al., 2007; Parker and Ingber, 2007). In vivo, it is the ECM which provides the basic physical, physicochemical and differentiative/instructional support system, thus contributing to functional differentiation and 3-D assembly of cells into tissues. Biomimetic scaffolds, which aspire to emulate many of the structural and functional characteristics of the ECM in vivo, are arguably one of the most important ‘engineering’ components in vitro, bioreactors being the other one. Biological scaffolds can truly be ‘engineered’, in terms of their constituent (bio) materials (which can be either synthetic, or natural, or blends of the two), their intrinsic morphology (e.g. porous, fibrous, amorphous/ hydrogel), as well as their capability to support important functional design requirements, such as tissue-specific 3-D assembly of cells into glandular structures, vascularization and innervation. Thus, since TE scaffolds aim at mimicking the multiple functions of the ECM, this area of research is a prime translational target for the biomedical application of natural-based polymers. There is a need in biomedical sciences for biomimetic scaffolds of biocompatible composition and of nanofibrous structure, which closely emulate the composition and structure of the natural ECM, and which can be implanted into the patient. Hence, considerable effort has been and is being invested into the development of (biodegradable) polymer scaffolding suitable for TE applications. Ideally, a candidate scaffolding should mimic the structural and functional profile of the materials found in the native ECM. One of the important criteria for selecting a (natural) polymer for use as a biomaterial is to match its mechanical properties and time of degradation to the needs of the application. Biodegradable polymers can be either natural or synthetic. In general, synthetic polymers offer certain advantages over natural materials in that they can be tailored to give a wider range of physicochemical properties and more predictable lot-to-lot uniformity than can materials from natural sources. Most commercially available biodegradable devices are polyesters composed of homopolymers or copolymers of glycolides and lactides (Middleton and Tipton, 1998). On the other hand, synthetic materials that have been used in attempts to meet the above criteria for scaffolds have largely failed to live up to expectations in the clinical setting (Matthews et al., 2002). The ability of cells to assemble into tissues and maintain tissue-specific functions critically depends on epigenetic factors, such as the unique cell/ tissue-specific microenvironment. Some of the major factors contributing to

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this unique microenvironment are cell-cell interactions and the organotypic ECM. Interactions between cells and ECM are crucial to cellular differentiation and in modulating or redirecting cell function. However, when cells are removed from their microenvironment and cultured in vitro, they typically dedifferentiate, thereby losing some of their normal in vivo behavior. A principal objective of TE, therefore, is to create an in vitro 3-D culture system that provides some of the essential factors in the microenvironment, which control and regulate cell function in vivo. As such, electrospun scaffolds from natural proteins could serve as excellent mimics of the tissue-specific ECM (Teo et al., 2006). In this chapter we will briefly introduce the principles of electrospinning and discuss electrospun TE scaffolds made mainly of natural animal proteins (mammals and spider). In addition, after a short discourse about using blends of synthetic and natural polymers as well as complex ECM protein mixtures, we will focus on the novel use of ‘green’ alimentary plant proteins for TE purposes. This is, in our view, a surprising twist to the concept of using lowcost, renewable resources for high-tech applications.

17.2

The electrospinning process

Electrospinning, originally invented in the 1930s in the textile industry (Formhals, 1934), is a convenient technique for producing non-woven fabrics with fiber sizes ranging from < 100 nanometers to tens of micrometers (Doshi and Reneker, 1995; Reneker and Chun, 1996; Frenot and Chronakis, 2003; Li and Xia, 2004). Antonin Formhals, the original inventor of the technology, demonstrated in 1934 that an electrostatic force could be used to produce polymer filaments. In the basic process, a polymeric melt or solution is exposed through a nozzle to an external electric field, characterized by very high voltage (10s of kV) and very low currents, thereby charging a reservoir of polymer fluid and accelerating a fluid jet through the electric field gradient toward a grounded target or collector (Ramakrishna et al., 2005). As the conical jet (Taylor cone) of polymer fluid propagates through the air, the solvent evaporates and a non-woven mat of submicrometerdiameter fibers is produced on the collector (Fig. 17.1a). The ability of the polymer to form chain entanglements in solution will determine whether or not fibers form on the collector (Fig. 17.1b) (Shenoy et al., 2005a). Liquid jet stabilization and the subsequent formation of fibers are attributed to molecular phenomena including physical entanglements and thermoreversible junctions (Shenoy et al., 2005b). Inter-chain hydrogen bonding that results from hydrophilic polymer-polymer interactions is another factor that can promote fiber formation (McKee et al., 2004). The concentration and/or viscosity of the polymer solution, surface tension (Magarvey et al., 1962), applied voltage, air gap distance, and delivery rate are critical

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17.1 (a) Schematic representation of the electrostatic process. In the basic process, a reservoir of polymer fluid is charged. If the electrostatic forces exceed the surface tension of the fluid, the fluid jet is accelerated through an electric field gradient toward a grounded target or collector. (b) In the electrospinning process, droplet fragmentation is limited by the presence of polymer (or protein) chain entanglements. Chain stretching is accompanied by the rapid evaporation of the solvent. Dry fibers accumulate on the surface of the collection plate in the form of a non-woven mat (Woerdeman et al., 2007b).

experimental processing variables which determine the shape and size of electrospun fibers (Katti et al., 2004; Ramakrishna et al., 2005). With prior knowledge of the entanglement molecular weight and weight-average molecular weight of a particular polymer in a ‘good solvent’, one can predict the critical polymer concentration needed to cross the transition between electrospraying and electrospinning (Shenoy et al., 2005a).

17.2.1 Effects of polymer concentration and viscosity In developing TE scaffolding from natural ECM proteins, we studied the effects of various electrospinning parameters to produce desirable electrospun fibers (Li et al., 2005). For economic reasons, most of our studies optimizing the electrospinning parameters of ECM-derived molecules were carried out with gelatin, rather than with collagen, and with alpha-elastin, rather than tropoelastin. However, all critical parameters used were also validated with these other proteins. An increase in the concentration of the polymer solution leads to a corresponding increase in the viscosity of the solution. Beads and beaded fibers, one of the possible undesired artifacts of electrospinning under suboptimal conditions, are less likely to be formed for the solutions with

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higher concentration and more viscous solutions. The size of the beads becomes larger and the average distance between beads on the fibers becomes longer as the concentration and viscosity increase. Meanwhile, the shape of the beads gradually changes from spherical to spindle-like (Fong et al., 1999), until, above a particular concentration that depends on a given combination of materials and solvents, beads disappear altogether and turn into fibers. Conversely, for a given molecular weight, an immediate consequence of reducing the solute concentration is a decrease in the ensuing fiber size. For example, by decreasing the concentration of gelatin in the solvent 1,1,1,3,3,3hexafluoro-2-propanol (HFP) from 8.3% to 2%, the average size of gelatin fibers was reduced from 485 ± 187 nm to 77 ± 41 nm (p<0.01) (Li et al., 2005). The reduction in gelatin concentration yielded the expected reduction in the average fiber size to below 100 nm; this reduction of fiber size, however, was accompanied by a significant formation of beads as seen in Fig. 17.2. Similar results were obtained for electrospun collagen fibers. In line with previous reports on other polymeric materials (Fong and Reneker, 1999), our data suggest that electrospinning of collagenous proteins at concentrations above 5% will yield smoother and more uniform fibers of

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17.2 SEM micrographs of electrospun gelatin fibers at different concentrations: (a) 3%; (b) 5%; (c) 8.3%.

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several hundred nanometers in diameter. By contrast, electrospinning at lower concentrations will result in smaller fibers, but with more beads if the entanglement consideration is not satisfied. One possible solution is to add a small quantity of a high molecular weight polymer to the solution. Shenoy et al. (2005a) have shown that the lower number of chains per unit volume directly translates into lower fiber diameters, since high molecular weight polymers have higher chain extensibilities than low molecular weight polymers. The diameter of electrospun α-elastin fibers also depended on the concentration of the solution; the dose response curve, however, was significantly right-shifted. As seen in the scanning electron microscopy (SEM) micrographs in Fig. 17.3, electrospinning of α-elastin at 10% yielded large beads and fragmented fibers; at 15%, there were no more beads, but the fibers were still fragmented. When the concentration of α-elastin was raised to 20%, the electrospun fibers were continuous and more uniform. These results are in line with the parametric investigations of optimizing the generation of thin protein films from silk-like polymer and fibronectin (Buchko et al., 1999). Generally, α-elastin fibers were larger in diameter (2–8 µm) and

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17.3 SEM micrographs of electrospun elastin fibers at different concentrations: (a) 10%; (b) 15%; (c) 20%.

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more ‘elastic’ than the ‘straight’ gelatin and collagen fibers (with diameters of 400–800 nm).

17.2.2 Effects of electrospinning voltage and air gap distance Electrospinning voltage and air gap distance are two other important parameters governing the uniformity of electrospun protein fibers. Keeping a fixed delivery rate of 5 ml/h, two voltages of 10 kV and 20 kV were chosen to study the effects of varying the air gap distance between 5 cm to 40 cm. As seen in Fig. 17.4, the smallest fiber size was obtained at an optimized ratio of applied 4.0

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17.4 Effects of electrospinning voltage and air gap distance: (a) Sizes of electrospun gelatin fibers vs. ratio of applied voltage to gap distance; (b) Ratios of standard deviation to average fiber size vs. ratios of applied voltage to air gap distance.

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voltage to air gap distance of 0.667 kV/cm (Fig. 17.4a). Spinning at 20 kV apparently resulted in smaller fibers than at 10 kV (p < 0.01). However, for the same ratio of applied voltage to air gap distance, electrospun fibers produced at 20 kV were much less uniform than those at 10 kV, as inferred from the larger standard deviation (Fig. 17.4b). Thus, a trade-off exists between the uniformity of bead-free continuous fibers vs. smaller fiber size. Deitzel et al. (2001) reported that with increasing voltage, the resulting poly-(ethylene oxide) (PEO) nanofibers became rougher and contained more beads. Although we did not measure the roughness of the electrospun protein fibers and density of the beads, we surmise that Deitzel’s observation also holds for fibers spun from biological polymers, such as the ones used in our study.

17.2.3 Effects of solution delivery rate The effects of the solution delivery rate on the size and the periodicity of biopolymeric protein fibers were investigated for gelatin and elastin fibers. Three of the electrospinning parameters were kept constant: concentration (8.3% for gelatin and collagen, 20% for alpha-elastin and tropoelastin), applied voltage (10 kV), and air gap distance (15 cm). The delivery rates of the protein solutions were systematically varied between 1 ml/h – 8 ml/h in order to optimize the conditions that yield bead-free and uniform fibers. As shown in Fig. 17.5, increasing the delivery rate from 1 ml/h to 3 ml/h yielded a relatively small but significant change by ~25% (p < 0.01) in fiber diameter from 431 ± 105 nm to 533 ± 119 nm and from 349 ± 97 nm to 460 ± 148 nm for collagen and gelatin, respectively. A further increase in the rate of delivery from 3 ml/h to 8 ml/h did not significantly change the mean diameters of either collagen or gelatin fibers (p > 0.05). By contrast, for α-elastin and even more so for tropoelastin, which intrinsically yield wider fibers than collagen or gelatin, the widths of the ensuing fibers were strongly affected by the delivery rate. With an increase in delivery rate from 1 ml/h to 8 ml/ h, the mean fiber width increased about seven-fold, from 0.6 ± 0.1 µm to 3.6 ± 0.7 µm (α-elastin) and from 1.4 ± 0.3 µm to 7.4 ± 2.3 µm (tropoelastin), respectively (p < 0.01). These values are also comparable to the sizes of electrospun bovine ligamentum nuchae elastin fibers of 1.1 ± 0.7 µm (Boland et al., 2004). As shown in the (autofluorescence) micrographs in Fig. 17.6, α-elastin and tropoelastin fibers electrospun at a delivery rate of 1.5 ml/h significantly differed from gelatin and collagen fibers, in that α-elastin and tropoelastin fibers attained an elastic, wavy pattern, while collagen and gelatin fibers were mostly straight. However, when electrospun at a delivery rate of less than 1.5 ml/h, α-elastin and tropoelastin did not show a wave-like pattern but appeared coiled, while collagen/gelatin fibers were mostly straight. An increase in the delivery rate did not visibly affect the topology of gelatin and

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17.6 Fluorescent images of electrospun fibers. (a) gelatin (b) elastin. Original magnification: 100x.

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collagen fibers. By contrast, the patterns of α-elastin and tropoelastin fibers changed greatly. Shown in Fig. 17.7 are SEM micrographs of 20% α-elastin fibers electrospun at different delivery rates. Upon increasing the delivery rate beyond 3 ml/h these fibers acquired a spring-like wavy pattern. At higher delivery rates (5 and 8 ml/h) individual fibers appeared coiled around a single straight axis. Tropoelastin fibers displayed patterns similar to α-elastin. In summary, our studies on natural ECM polymers yielded results very similar to those reported previously for synthetic polymers (Katti et al., 2004; Mo et al., 2004) in terms of the parameters (e.g. solute concentration (most important), voltage, distance, and delivery rate) that are critical in fine-tuning the fiber shape and size of scaffolds electrospun from natural ECM polymers.

17.3

Electrospinning natural animal polymers

17.3.1 Gelatin, collagens, elastin and tropoelastin Type I collagen and elastin are two of the key structural proteins found in the extracellular matrices of many tissues (Toshima et al., 2004; Ntayi et al.,

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17.7 SEM micrographs of elastin fibers electrospun at different delivery rates: (a) 1 ml/h; (b) 3 ml/h; (c) 5 ml/h; (d) 8 ml/h.

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2004). These proteins are important modulators of the physical properties (stiffness/ elasticity) of any engineered scaffolds, affecting cellular attachment, growth and responses to mechanical stimuli (Lu et al., 2004; Buijtenhuijs et al., 2004; Kim and Mooney, 2000). Interstitial collagens and elastin provide instructive/differentiative cues, interacting with cells through integrins as well as through non-integrin receptors (Hinek, 1996; Rodgers and Weiss, 2005; Leitinger and Hohenester, 2007). Of the more than two dozen members of the collagen family that are currently known, only a few have been electrospun, most frequently type I collagen, the ubiquitous structural collagen, found in nearly all tissues. To the best of our information, Chaikof and coworkers were the first to electrospin type I collagen scaffolds for wound dressings (Huang et al., 2001a; Huang et al., 2001b). Shortly thereafter Bowlin and coworkers described electrospinning of type I collagen and elastin fibers for preliminary vascular TE (Matthews et al., 2002; Boland et al., 2004). For a recent summary of the many facets of electrospun type I collagen see the excellent overview by Barnes et al. (2007) and Murugan and Ramakrishna (2006). By contrast, only few papers have reported electrospinning of type II collagen, the predominant collagenous component of articular cartilage (Shields et al., 2004; Barnes et al., 2007). To date only one paper (Matthews et al., 2002) briefly mentioned electrospinning type III collagen, the second most abundant ubiquitous collagen type that is found in numerous elastic tissues, e.g. lung, skin and vasculature. To the best of our knowledge, and with the exception of data shown for the first time later (see Section 17.3.2), no studies have been reported so far attempting to electrospin other types of collagen, including the all-important type IV collagen found in the basement membranes of all epithelia and endothelia. Thus, if one wants to study the role of tissue-specific collagens in tissue assembly in health and disease, there is clearly a need for generating biomimetic scaffolds comprising various types of collagens, and not just type I collagen. In addition to differences in their diameters (see above), there are significant differences in the topology of the fibers spun from collagenous proteins and from elastin/tropoelastin. As seen in the SEM and atomic force microscopy (AFM) micrographs in Fig 17.8, both gelatin and collagen fibers appear uniformly round when electrospun at low delivery rates, around 1 ml/h. By contrast, under the same experimental conditions, α-elastin as well as tropoelastin fibers appear wider and flatter, shaped like ribbons, their size resembling that of naturally occurring elastin fibers (Leppert and Yu, 1991). Analysis by AFM indicates that the elastin/tropoelastin ribbons exhibit a symmetric increase in their thickness at the edges (Fig. 17.8d). In situ, elastin has a wave-like periodic appearance in the larger elastic arteries (Birk et al., 1991). Interestingly, the innate elastic properties of α-elastin and tropoelastin are retained upon electrospinning. Fibers made of

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µm 4

2

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17.8 SEM and AFM micrographs of electrospun fibers: (a, c) collagen fibers; (b, d) tropoelastin fibers.

α-elastin and tropoelastin attained an elastic, wavy pattern, while collagen and gelatin fibers were mostly straight (see Fig. 17.6). As shown in Fig. 17.9, the periodicity of the waves of tropoelastin fibers depended on the delivery rate. When electrospun at 1 ml/h, the periodicity was the smallest (71 ± 28 µm). With increasing delivery rate, the periodicity gradually increased to 105 ± 22 µm (at 3 ml/h) to a maximum of 126 ± 25 µm (at 7 ml/h). To examine the tensile properties, we performed microtensile tests on electrospun fiber sheets. The average tensile moduli are listed in Table 17.1. Electrospun collagen fibers have a lower tensile modulus than gelatin fibers, but have similar tensile strength (8-12 MPa) and ultimate elongation (0.080.1), respectively. Moreover, elastin fibers are much more brittle than gelatin and collagen, or even tropoelastin fibers: the tensile strength of elastin fibers is only about 1.6 MPa and the ultimate elongation is about 0.01. Tropoelastin is more elastic than elastin, and also more elastic than gelatin and collagen fibers: its ultimate elongation reaches 0.15 and the tensile strength reaches almost 13 MPa. By comparing their elastic properties, we surmise that tropoelastin is advantageous over elastin for fabricating engineered scaffolds which could mimic the in vivo ECM environment. For TE applications, i.e. to realistically emulate characteristics of natural tissues, the mechanical

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y = 26.631 Ln (x) + 70.378 R2 = 0.9646

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Periodicity (µm)

140

*

* *

120

*

100 80 60 40 20 0 0

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4 5 6 Delivery rate (ml/h)

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10

* p < 0.05, values are significantly different from the first one.

17.9 Periodicities of electrospun tropoelastin fibers at different delivery rates. Table 17.1 Tensile moduli of electrospun protein fibers (MPa) (n = 3) Gelatin

Collagen

Elastin

Tropoelastin

426 ± 39

262 ± 18

184 ± 98

289a

Note: aThe limited availability of the recombinant material precluded extensive testing of this material; the value listed is the average of two independent tests with similar results.

properties of gelatin fiber scaffolds would have to be enhanced, for example by co-spinning with other synthetic polymers such as polyurethane, PLGA or by the addition of carbon nanotubes (Weisenberger et al., 2003).

17.3.2 Complex ECM protein blends (MatrigelTM) To date, only single defined proteins, or, as discussed below, blends of well defined natural and synthetic polymers have been electrospun for scaffolding in TE. However, the ECM that these scaffolds arguably try to emulate is much more complex, comprising a plethora of functional and structural macromolecules. In addition, the ECM also harbors numerous trophic factors which are secreted and/or deposited by the cells in a tissue- or diseasespecific manner and may become bio-available in a tightly coordinated fashion. The microscopic/nanoscale structure of the ECM molecules, the spatiotemporal availability of the instructive cues contained in these molecules and their degradation products, and their interplay are important for orchestrating appropriate cell functions. As discussed above, the many differentiative cues

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of the ECM proteins are largely absent in synthetic scaffolds. Some ECM protein-derived scaffolds can act locally as biomimetics, facilitating tissue repair without additional inclusion of exogenous growth/differentiation factors or cells (for a review, see Badylak, 2007). Hence, a major challenge for TE is to generate scaffolds which are sufficiently complex in mimicking the differentiative/instructive functions of the native ECM and are not immunogenic. MatrigelTM (MG) is a readily available, complex ECM extract. It is isolated from the murine Engelbreth-Holm-Swarm (EHS) sarcoma, and contains a complex mixture of basement membrane proteins, mainly laminin and collagen IV as well as heparan sulfate proteoglycans, entactin and nidogen. In addition, MG contains a number of bioactive molecules and peptide growth factors, such as epidermal growth factor (EGF), transforming growth factors – βs (TGF-βs), platelet-derived growth factor (PDGF) and many others (for a recent review, see Kleinman and Martin, 2005). Unlike synthetic scaffolds, MG provides a more natural, biocompatible environment to cells and promotes the growth, tissue-specific morphogenesis, and differentiation of stem cells (Chen et al., 2007) as well as differentiated cells that are otherwise difficult to grow/maintain in a tissue-specific fashion in vitro, such as neurons, hepatocytes, sertoli cells, hair follicles, thyroid cells and epithelial cells (Mondrinos et al., 2006). MatrigelTM is liquid at 4°C but forms a semi-solid, viscous hydrogel at 37°C. Thus, in its present incarnation, it may be of limited usefulness as a scaffold for TE purposes. We hypothesized that we could generate more complex, fibrous scaffolds by electrospinning MG, which would retain (at least in part) its superior bioactivity. Upon extraction from the EHS tumor, MG was lyophilized and then solubilized in HFP at a concentration of 20% (w/v) and electrospun following the protocols developed for other natural proteins (see above). Shown in Fig. 17.10 are SEM micrographs of electrospun MG fibers before and after crosslinking with 1-ethyl-3-[3-dimethylaminopropyl]carbodiimide hydrochloride (EDC). Under optimized conditions, the size of MG fibers (a)

(b)

17.10 SEM micrographs of electrospun Matrigel fibers: (a) before; and (b) after crosslinking with EDC.

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was about 1.16 ± 0.33 µm (n = 30). In our previous experiments, the diameter of collagen fibers was 349 ± 97 nm (n = 30) and elastin fibers was 605 ± 102 nm (n = 30) under similar electrospinning conditions (Li et al., 2005). One reason for the larger fiber size may be that the viscosity of a 20% MG solution is higher than that of a 20% elastin solution. In tensile tests, the maximum strain of MG fibers was 0.52 ± 0.10 (n = 5) and the secondary modulus was 56.0 ± 12.2 MPa (n = 5). In comparison to the previously (Li et al., 2005) reported tensile properties of electrospun collagen (max strain: ~0.1; modulus: ~10 MPa), elastin (max strain: ~0.01; modulus: ~1.6 MPa) and tropoelastin (max strain: ~0.15; modulus: ~13 MPa), our data clearly indicate that MG fibers represent a novel natural material with physicochemical properties unique and dissimilar from those of each of the individual materials, i.e. collagen, elastin, or tropoelastin, and others.

17.3.3 Fibrinogen Fibrinogen is an essential plasma protein, the precursor to fibrin during clot formation. Fibrin has been used for a long time as a permissive provisional matrix for TE. The physical properties of fibrin hydrogel matrices, and, with that, their usefulness as provisional TE scaffolds, can be easily modulated by adjusting the concentrations of fibrinogen, calcium and thrombin (Ye et al., 2000; Linnes et al., 2007). More recent studies indicated that mammalian cells, notably epithelial cells, secrete and process fibrinogen independent of its role as a plasma protein, and that this event may be part of natural healing phenomena (Perrio et al., 2007). Thus, fibrin glue is not only fortuitously an excellent building block for biomimetic scaffolds, but an important ingredient in the repair process. Hence, we and others have attempted to electrospin fibrin scaffolds, albeit to date with limited success mainly because of technical obstacles in the timing of the thrombin-induced fibrinogen cleavage and fibrin polymerization during the spinning process. Rather, the Bowlin group (McManus et al., 2006; McManus et al., 2007) as well as our group (see Fig. 17.11) have electrospun fibrinogen to yield biomimetic scaffolds with fiber sizes in the submicron range. Shown in Fig. 17.11 is an SEM micrograph of electrospun, aligned fibrinogen fibers, prepared out of a 10% (w/v) fibrinogen solution in HFP. Following establishment of the feasibility of this approach, Sell et al. (2007) recently developed a new methodology for assessing fiber diameters in crosslinked fibrinogen matrices by measuring scaffold porosity.

17.4

Electrospinning blends of synthetic and natural polymers

Nanofibrous scaffolds electrospun from natural ECM protein have shown improved cellular responsiveness, mainly because of their physicochemical

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17.11 SEM micrograph of electrospun fibrinogen fibers.

similarity to the native ECM. However, due to the need for chemical crosslinking and poor mechanical properties of electrospun natural proteins, incorporation of synthetic polymers or of other materials (e.g. carbon nanotubes) might be needed to enhance the mechanical properties of fibrous scaffolds or to produce novel materials with unique properties for custom-tailored applications in TE.

17.4.1 Gelatin/collagen, elastin and PLGA blends Buttafoco and coworkers (2005, 2006) were the first to co-electrospin a binary blend of collagen and elastin from aqueous solutions of these natural proteins. However, in order to obtain homogeneous and continuous nanofibers, the authors had to add poly(ethylene oxide) (PEO 900 kDa) and NaCl. The addition of PEO stabilized the jet by increasing the viscosity of the blend, while the addition of low concentrations of NaCl produced uniform fibers. Chemical cross-linking is necessary to stabilize these nanofibers in aqueous environments, although both PEO and NaCl completely evanesce after crosslinking, as assessed by differential scanning calorimetry and SEM. Schnell et al. (2007) reported that a binary blend of natural proteins (e.g. collagen or gelatin) with synthetic polymers, such as poly(lactide-co-glycolide) (PLGA) or poly(ε-caprolactone) (PCL) enhances the mechanical properties of the resultant fibrous scaffolds. However, all these binary blends still need to be cross-linked in order to maintain fiber integrity in an aqueous solution. We recently generated a new class of biohybrid scaffolds by coelectrospinning tertiary blends of PLGA (90/10), gelatin, and α-elastin, and refer to this new material as PGE (Li et al. 2006b; Han et al., manuscript in preparation). After PLGA, gelatin and α-elastin were dissolved in HFP at optimized concentrations of 10%, 8% and 20% (w/v), respectively; tertiary PGE blends were mixed at different volume ratios. Electrospinning was carried out as described above (Li et al., 2005; Li et al., 2006). Surprisingly, the resulting fibers were much smaller than each of the individual components

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and homogeneous without evidence of phase separation (Fig. 17.12a). More importantly, the fibrous PGE scaffolds swelled upon hydration, as indicated by an increase in average fiber diameter from 380 nm to 856 nm (Fig. 17.12b), and were stable in aqueous environments without need for chemical cross-linking. The hydrated fibers also showed a reduction in Young’s modulus from 141 MPa to 43 MPa. In terms of biological properties, PGE scaffolds supported H9c2 rat cardiac myoblast attachment and proliferation under static conditions as shown in Fig. 17.13. Hydrated PGE scaffolds resemble opaque hydrogels, supporting the notion that the diverse materials in the homogeneous PGE blend are arranged in such a way that PLGA serves as the backbone while the water soluble gelatin/ elastin face the aqueous surface. The morphology of individual fibers after hydration and swelling was influenced by the different volume fractions of PLGA and gelatin in the blends: fibers with a high relative content of PLGA

(a)

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17.12 SEM micrographs of PGE fibers: (a) dry fibers; (b) hydrated fibers soaked in cell culture medium for 36 h and then baked on a hot plate. Original magnification: 5000x.

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17.13 Fluorescent and SEM images of confluent H9c2 myoblasts on PGE fiber-coated glass coverslips eight days post-seeding. Staining for nuclei-bisbenzimide and actin cytoskeleton-phalloidin. Original magnifications: 400x for (a) and 2500x for (b).

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and/or low relative content of gelatin swell less than the fibers of low relative content of PLGA and/or high relative content of gelatin. We conclude that gelatin tends to flatten the hydrated fibers, while PLGA tends to constrain the fibers and thus stabilize the whole construct. Therefore, these two materials play a synergistic role in modulating the biochemical and mechanical properties of the PGE constructs. More recently, we systematically varied the ratios of PLGA and gelatin in order to assess the contribution of each of individual material component to modulating the physicochemical and biological properties of the PGE scaffolds (Han et al., manuscript in preparation). Fibrous scaffolds were fabricated by parametrically varying the volume ratios of the three individual components. Preliminary data indicate that a ratio of P:G:E (3:2:1) produced the smallest average fiber size (~300 nm) and its hydrated mats exhibited the highest Young’s modulus (0.770 MPa) and tensile strength (0.130 MPa). These data further suggest that the elasticity of optimized PGE scaffolds is in the range of that of natural blood vessels. The relative ratio of PLGA appears to be the single most critical factor controlling the mechanical properties of the PGE fibrous scaffolds; there is little relationship between the fiber size and mechanical properties of the PGE scaffolds. Hypothesizing that suitable blends of natural and synthetic polymers might yield a potentially anti-thrombogenic scaffold for application as small diameter vascular grafts, we seeded endothelial cells (EA.hy926) and bovine aortic smooth muscle cells (BASM) onto PGE scaffolds. As shown by routine histological analysis (Fig. 17.14), the endothelial cells formed monolayers on the PGE scaffolds while the smooth muscle cells penetrated into the scaffold interior and formed multiple layers in the scaffolds, reminiscent of their organization in situ. These encouraging preliminary findings suggest

(a)

(b)

17.14 Histological analysis (H&E staining) of cell-seeded PGE scaffolds: (a) populated with a confluent monolayer of endothelial cells (EA.hy926 cells); (b) bovine aortic smooth muscle cells. Magnifications are 200x for (a) and 400x for (b).

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that the PGE fibrous scaffolds support cell-type specific localization of cells seeded onto the grafts and bode well for the notion that in the not too distant future, PGE might be custom-tailored as potential anti-thrombogenic scaffolds for vascular grafts. In line with our studies, Stitzel et al. (2006) and Lee et al. (2007) coelectrospun a tertiary blend of collagen, PLGA and elastin blends as potential nanofibrous scaffolds for vascular graft application. Lee et al. also compared the average fiber sizes and mechanical properties of vascular grafts coelectrospun from blends of collagen and elastin with PLGA, poly(L-lactide) (PLLA), PCL and poly (D,L-lactide-co-ε-caprolactone) (PLCL). Their data showed that the scaffold electrospun from PLGA blends produced the smallest average fiber size (~ 400 nm) while the one from PLLA had the most natural vessel-like mechanical properties, with a Young’s modulus 2.08 MPa and tensile strength 0.83 MPa. Due to the high cost of collagen, gelatin might provide a feasible alternative in co-electrospinning natural and synthetic blend materials. Schnell et al. (2007) co-electrospun another synthetic polymer PCL with collagen for neural TE. Electrospun collagen/PCL nanofibrous scaffolds provided improved glial cell guidance/glial cell formation and migration, neurite orientation and axonal growth as compared to pure electrospun PCL fibers. Another study (Zhang et al., 2005) compared electrospun collagencoated PCL nanofibers with roughly collagen-coated PCL matrix (obtained by soaking PCL in collagen solution) and showed that the electrospun fibers could be used as functional biomimetic matrices with excellent cell-scaffold integration.

17.4.2 Polyaniline (PANi)-contained gelatin nanofibers Electrical current/activity seems to be beneficial for differentiative and regenerative processes (Bidez, 2006; Song, 2007). One possible approach to translate these observations into the realm of TE is to endow biomimetic scaffolds with an added degree of ‘intelligence’, by co-electrospinning natural and conductive polymers such as polyaniline (PANi) or polypyrrole. We hypothesized that such electrical stimulation via the scaffold would modulate pivotal cell functions, such as cell attachment, migration, proliferation and differentiation. Based on our previous experiences with conductive PANi substrates inducing differentiation of mouse embryonic stem cells and PC12 cells into the neuronal and cardiac lineages respectively (Bidez, 2006; Guo et al., 2007), we co-electrospun a novel blend of conductive camphorsulfonic acid-doped emeraldine PANi (C-PANi) with gelatin to generate conductive TE scaffolds (Li et al., 2006a). The addition of PANi to gelatin produced homogeneous fibers without phase separation. As before, with increasing concentration of the synthetic polymer, in this case C-PANi, the average

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fiber size was drastically reduced from approximately 800 nm to 60 nm and the tensile modulus of the scaffolds was increased from approximately 499 MPa to 1384 MPa. To evaluate the biological properties of the resultant CPANi-gelatin blend fibrous scaffold, we cultured H9c2 rat cardiac myoblasts on these scaffolds, demonstrating that this blend of fibrous scaffold supported cellular attachment, migration and proliferation. Electroactive fibrous scaffolds might be of particular value in cardiac TE, in cases where native tissue has lost at least part of its function due to myocardial infarction or other injury.

17.4.3 Silk fibroin blends Silk fibroin (SF) is the natural, purified fiber produced by silkworms. It was one of the first natural animal proteins to have been electrospun (Jin et al., 2002), and has since been extensively investigated as one of the most promising candidate biomaterials for its good biocompatibility, biodegradability and minimal inflammatory reaction. Raw silk contains the fibrous protein, fibroin, as the thread core while glue-like proteins named sericin surround the core to cement it together. Kaplan and coworkers pioneered the electrospinning of silk and improved the spinnability and processability of SF by incorporating PEO and avoiding the conformational transitions during solubilization, similar to the case of the elastin/gelatin blend discussed before. Various compositions of SF/PEO aqueous blends have been electrospun and the resultant fibers were uniform with composition reflective of the solution concentration and less than 800 nm in size (Jin et al., 2002). More recently, Ayutsede et al. (2006) introduced electrospun nanocomposite fibers of worm silk and single wall carbon nanotubes (SWNT) by dispersing SWNTs in SF solutions, with cohesive forces mainly driven by steric and hydrophobic effects. The resultant multifunctional, strong and tough fibers showed a significant increase in the

(a)

(b)

5 µm

1 µm

17.15 SEM micrographs of 1% SWNT reinforced fibers: (a) aligned; and (b) random with a weblike structure.

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initial Young’s modulus in the range of 110~460% (Fig. 17.15). What remains to be improved is the homogeneous distribution of SWNTs in the electrospun fibers, before potential application of the novel nanocomposite fibers as TE scaffolds. Because of chitin’s superior hydrophilicity and degradability, Park et al. (2006) also incorporated chitin in SF blends in order to increase the hydrophilicity, degradability and biofunction of the resultant blend silk/chitin fibers. The average fiber diameter of chitin/SF blend nanofibers was reduced from 1260 nm (SF) to 340-920 nm. In this blended fiber, SF was dimensionally stabilized by water vapor treatment, a technology used to induce SF crystallization. A chitin/SF blend of 75%/25% produced biomimetic threedimensional structure and supported excellent cell attachment and spreading. Therefore, the new family of silk-based fibrous scaffolds might be of great potential in TE applications.

17.5

Electrospinning novel natural ‘green’ plant polymers for tissue engineering

In the interest of environmental friendliness, low cost, and high availability of raw material, ‘green’ proteins derived from renewable plants such as soybean, corn and wheat have been investigated for industrial uses such as textiles, films, adhesives and plastics (Domenek et al., 2004; Brown, 2005; Subramanian and Sampath, 2007). In the TE arena, these same materials are being explored as potential scaffold biomaterials for the cost and availability advantages as well as the avoidance of immunogenic reactions and disease transfer risks associated with animal-derived proteins. Reis and colleagues are using, for example, corn starch and other plant products as base material for biodegradable biomimetic scaffolds and for drug delivery (for a recent review see Malafaya et al., 2007). Those studies will be discussed in greater detail elsewhere in this book. In this section, we highlight the recent progress in electrospinning fibrous scaffolds from three alimentary proteins: wheat gluten, corn zein and soy protein.

17.5.1 Soy proteins Soy protein is an abundant globular protein that comprises two main classes: 10% of the proteins consist of albumins, which can be extracted by water. The other 90% is made up of globulins, which can be extracted by dilute salt solutions. The globulin proteins can be subdivided further according to their sedimentation rates when dissolved in pH 7.6, 0.5 M ionic strength buffer into the following four fractions: 2S (15%), 7S (34%), 11S (41.9%) and 15S (9.1%) (Nielsen, 1985; Wool and Sun, 2005). Soy protein is commercially purified as an isolate of purity greater than 90% and consists of 7S and 11S

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proteins. The 11S fraction is pure glycinin, while the 7S fraction consists of mostly β-conglycinin, as well as small quantities of γ-conglycinin, lipoxygenases, α-amylases, and hemagglutinins (Nielsen, 1985). Glycinin comprises one basic polypeptide and one acidic polypeptide linked together by a disulfide bridge. At pH 7.6 and at room temperature, glycinin forms hexameric complexes with a molecular weight of around 360 kDa. The isoelectric point of glycinin is 4.9 (Koshiyama, 1983). β-conglycinin, on the other hand, is a trimeric glycoprotein and consists of three different polypeptide subunits (α′, α, β) with molecular weights ranging from 57–72 kDa, 57–68 kDa, and 45–52 kDa, respectively (Yamauchi, 1991). The isoelectric point of β-conglycinin is 4.64 (Koshiyama, 1983). Soy protein has been extensively studied for its intrinsic properties as well as its film-forming and gelation abilities (Jiang et al., 2007b; Mauri and Añon, 2006; Maltais et al., 2007), but very little work has been documented so far on its use in fiber form for TE scaffolding. So far, it has been difficult to obtain consistent electrospun fibers, when using soy protein isolate (SPI) alone. Rather, in our preliminary efforts we tended to observe discrete agglomerations exhibiting electrostatic spraying behavior. Thus, as discussed above, addition of a small percentage of a high molecular weight carrier polymer will increase chain entanglements, and enable the formation of continuous fibers. Early work by Zhang et al. (1999) used a bi-component wet fiber spinning apparatus and a rather complex manufacturing procedure, involving protein denaturation in a mixture of sodium hydroxide, urea and sodium disulfite, to produce soy protein-poly(vinyl alcohol, PVA) fibers with a core-sheath structure. In order to simplify the process the same group subsequently generated wet spun PVA/soy protein blend fibers, using thermal rather than alkali denaturation (Zhang et al., 2003). A PVA/soy protein ratio of 90/10 yielded fibers with a much greater strength (145 ± 10 MPa) than those from a 20/80 ratio (48 ± 6 MPa) when crosslinked with glutaraldehyde. However, crosslinked fibers with the greater relative amount of protein had a much higher value for elongation at break (52 ± 6%) than the 90/10 ratio (12 ± 1%). Our own initial attempts at electrospinning pure SPI were also carried out in dilute sodium hydroxide and yielded only electrostatic spraying behavior over a wide range of solution concentrations and electrospinning parameters tested. Addition of small amounts (less than 1%) of high molecular weight PEO to the SPI solution resulted in fiber formation. However, due to the instability of the protein solution in sodium hydroxide (the protein is prone to hydrolysis), we switched to the use of an organic solvent, HFP. In this solvent we consistently obtained ribbon-like fibers from SPI/PEO blend solutions (Fig. 17.16). Currently we are characterizing these fibers and investigating their interactions with cells. Using human dermal fibroblasts

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our in vitro cell culture experiments have shown that the fibrous scaffolds electrospun from both SPI/PEO (and corn zein) support cellular growth and proliferation, as well as the retention of normal cellular morphology. Further insight into specific interactions between plant protein-derived bioactive peptides and cultured human cells would allow us to better tailor these alimentary protein-derived scaffolds for individual TE requirements.

17.5.2 Corn zein Zein is the primary protein in corn and is one of the main co-products of the bio-ethanol industry. As others have noted, to make ethanol production economically feasible with less reliance on government subsidy, it is critical that the co-products of the industry be better utilized (Shukla et al., 2000). Zein, a prolamine that serves as the major storage protein in the endosperm of corn, is insoluble in water but can be solubilized in alcohol, urea or alkali solution. Zein is isolated directly from corn gluten meal as an industrial polymer. Zein grade F4000 (Freeman Industries LLC, Tuckahoe, New York) comprises 91.5% protein, 5.0% fat, 0.04% fiber and 0.05% ash. Moisture content can vary between 3.5% and 6.0% (Selling et al., 2004). Selling et al. used polyacrylamide gel electrophoresis (PAGE) to define zein aggregation and found significant protein bands at 21 kDa, with a main band at 18 kDa, and minor bands at 12, 46, 41 kDa. In two recent studies, ethanol has been used as a solvent for zein solutions in electrospinning (Yao et al., 2007a; Torres-Giner et al., 2008). Yao et al. investigated ethanol/water ratios of 70:30, 80:20 and 90:10 to determine the effect of ethanol concentration on fiber size and morphology and found that the morphology was similar among the ratios, but that fibers spun from the 70:30 ratio solvent were ‘softer and more lustrous’. They subsequently used

(a)

(b)

17.16 SEM micrographs of electrospun fibers from a blend of soy protein isolate and PEO. The fibers display a ribbon-like morphology characteristic of volatile solvents, and a heterogeneous size distribution typical of fibers electrospun from a blend of natural and synthetic polymers. Scale bar: (a) 20 µm; (b) 5 µm.

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this as their optimized ratio and found a biphasic effect of the zein concentration on the physical properties of electrospun fibers. At concentrations below 30% (w/v) the beaded fibers were too fragile to be handled, while the strongest fibers (1.70 MPa) came from solutions containing 40% (w/v) zein. Notably, both higher (50%) as well as lower (30%) concentration thresholds yielded weaker fibers (0.71 MPa and 0.57 MPa, respectively). This may be attributed to increasing fiber diameters with increasing zein concentrations, similar to the case of electrospinning animal proteins. The fiber diameter increased drastically by approximately four-fold (from ~ 1.5 µm at 30% to ~ 6 µm at 50%), resulting in greater overlapping surface area and possibly a much lower degree of intertwining. The morphology and thermal properties of zein nanofibers have been thoroughly characterized by Torres-Giner et al. (2008) using various concentrations of ethanol up to 96% (w/v), ethanol/acetic acid, and ethanol/ sodium hydroxide as solvents: acidified mixtures yielded flattened fibers with a higher glass transition temperature than fibers spun from aqueous ethanol solutions. Alkaline mixtures yielded solutions with lower viscosities at the same concentrations, and electrostatic spraying rather than electrospinning behavior. Typical of electrospun polymers in general, zein fiber diameters increased with increasing solution concentration and applied voltage, in a range of 25 to 50% and 7 to 17 kV respectively, and decreased with increasing tip-tocollector distance (Torres-Giner et al., 2008). Similarly, there was a pronounced effect of both the solution pH and its viscosity on the quality of the ensuing fibers, as shown in Table 17.2. Also in this study, zein fiber diameter was relatively constant at an average of less than 200 nm at ethanol content of solvent between 50 and 80%, but above 80% the diameter increased up to around 500 nm in 96% alcohol. This is most likely explained by the difference in net volume charge density of the solution jet. Increasing the amount of ethanol lowers the net volume charge density due to the accelerated solvent evaporation rate. Jiang et al. (2007) characterized fibers electrospun from zein solutions in N,N-dimethylformamide (DMF), noting average diameters of below 100 nm at a concentration of 400 mg/ml, increasing to around 400 nm at 600 mg/ml. Fibers were tubular and relatively uniform. However, in our own preliminary Table 17.2 The correlation between pH, solution viscosity and the resulting structure of electrospun zein fibers (Torres-Giner et al., 2008). pH

Viscosity (cP)

Structure

3.88 ± 0.42 5.97 ± 0.31 11.33 ± 0.36

242.80 ± 5.98 108.04 ± 7.63 36.24 ± 4.73

Flat fiber Tubular fiber Beads with nanocrystals

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attempts to replicate these results, the solution formed a gel at 400 mg/ml and could not be electrospun. Tubular, uniform fibers (Fig. 17.17) have been electrospun from solutions of corn zein dissolved in pure glacial acetic acid (Lin et al., manuscript in preparation). Zein solutions were fully dissolved within a few hours and remained stable over a period of several days: fibers spun from a fresh solution that had been stirring for only three hours were of the same size and morphology as those from a solution of the same concentration that had been stirring overnight. As with other polymers, solution concentration was the single most important determinant of fiber diameter, with ~159 nm at 35%, 240 nm at 40% and 351 nm at 45%.

17.5.3 Wheat gluten Gluten is the protein, starch, and lipid nutrient storage component of grain endosperm cells. The molecular-scale behavior of wheat gluten protein has been investigated extensively over recent decades (Wall, 1979; Redl et al., 1999; Singh et al., 1990; Kasarda, 1989). Commercial wheat gluten protein is highly complex and heterogeneous, in terms of both its molecular weight (ranging from 104–106 Da) and overall composition (comprising roughly 75% protein, 10% starch, 5% lipids, 5–10% water and <1% minerals). Both the high molecular weight glutenin polymers and low molecular weight gliadin monomers in wheat gluten exist as coiled or folded chains; the proteins are stabilized via disulfide bonds between cysteine residues. Mixing the gliadin and glutenin protein chains in solution or in a dough-like state promotes stretching of these molecules and simultaneously disrupts the weak bonds in

17.17 SEM micrographs of electrospun fibers from zein solutions in glacial acetic acid. Scale bar: 5 µm.

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the system. When chemical and physical bonds reform between the individual protein chains (e.g. by removing unbound solvent, or by allowing molecular relaxation to occur in the case of the dough), the disulfide bonds have the potential to form both intramolecular and intermolecular bonds between different protein molecules. Intermolecular bond formation via these disulfide bonds enables the gluten protein chains to achieve very high molecular weights. Wheat gluten, when mixed with water or another plasticizer, can be transformed into a bio-based polymer with thermoplastic properties. Standard petrochemical polymer molding processes can be used to form high-quality bio-based films and plastics derived from wheat gluten (Gennadios and Weller, 1990; Woerdeman et al., 2004). Under optimized conditions, fibers can be formed from wheat gluten (Woerdeman et al., 2005, 2007a, 2007b). The sulfhydryl/disulfide interchange reactions that occur among the glutenin polymers play a critical role in the fiber formation process. Wheat gluten was successfully electrospun using HFP and no additives. All electrospun wheat gluten mats were composed of highly heterogeneous flat ribbon-like fibers of varying widths ranging from tens of nanometers to over 3 microns. Polarized optical microscopy revealed the presence of a skin-core structure. This preliminary work also provided insight into the chemistry involved in making this protein amenable to electrospinning. Denatured wheat gluten, on the other hand, could not be electrospun into fibers. This may have been due to the stabilization of the unfolded state by an excess of sulfhydryl/ disulfide interactions. Protein denaturation can also lead to significant changes in solubility. Reddy and Yang (2007) demonstrated that wheat gluten can be processed into rather large fibers (~ 34 denier or ~ 96 µm) with mechanical properties similar to those of wool, stronger and more elastic than soy protein or zein and with weak alkaline and acid resistance comparable to PLA fibers at high temperatures. However, rather than by electrospinning, these fibers were produced by extrusion into a coagulation bath. Breaking tenacity of the fibers depended on solution concentration, aging time and temperature, with a narrow range of all three parameters for the highest breaking tenacity (ca. 120 MPa). Preliminary comparison of solely the mechanical properties of wheat gluten, as discussed above, with those of soy protein and zein, suggests that wheat gluten would be a more suitable raw material for TE scaffolds. However, processing conditions were different between the methods used to produce these data, and a fair comparison can only be made when electrospun fibers of the different alimentary proteins are tested under the same criteria. Most importantly, cellular responses to the different protein scaffolds must be characterized in parallel.

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17.5.4 Blends of synthetic and plant proteins Wheat gluten has been co-electrospun with poly(vinyl alcohol), PVOH, in order to generate electrospun scaffolds with improved mechanical properties (Woerdeman et al., 2007a,b). Micromechanical tests were conducted on fibrous mats electrospun from wheat gluten containing various amounts of PVOH: 0% (w/w) PVOH, 13% (w/w) PVOH, and 26% (w/w) PVOH, respectively (Fig. 17.18). Tensile strength (Fig. 17.18a) and percentage elongation-atbreak (Fig. 17.18b) demonstrate that gluten can be combined with a synthetic hydrophilic polymer to yield hybrid electrospun fibrous mats with significantly improved mechanical properties over those comprising only commercial wheat gluten. The tensile strength of the wheat gluten fibrous mats with 13% (w/w) PVOH was an order of magnitude higher than those without PVOH, and the tensile strength increased by another 100% upon increasing the percent of PVOH from 13% to 26% (Fig. 17.18a). More than a three-fold increase in the percentage elongation-at-break was achieved when the weightpercent of PVOH was increased from 13% to 26% (Fig. 17.18b). Although zein can be electrospun into nanofibers without any additives, blends of zein and other materials have also been explored (Yao et al., 2007b). Yao et al. produced ribbon-like composite fibers from a blend of zein and hyaluronic acid (HA), with poly(vinyl pyrrolidone), PVP, as a water-soluble synthetic additive to ‘facilitate the compatibility of zein and HA’. The mixture was electrospun in aqueous ethanol to avoid the toxicity of organic solvents or acids. As such, zein may be used as a complement to natural ECM proteins, and contribute its benefits in the form of mechanical property reinforcement. Thickness measurements for textiles are known to be complex (Gibson et al., 1999) and can be a source of error when incorporated into reported measurements of tensile stress. In this particular study, the thickness measurements were obtained without compacting the non-woven fibrous mats. Further, these tensile stress data do not account for changes in the fiber architecture that are likely to occur in the non-woven fibrous mats as a result of them being subjected to a mechanical load. In summary, alimentary ‘green’ proteins present a novel renewable resource for electrospinning scaffolds for TE purposes. This is one of those cases of significantly ‘added value’, by which an apparently inexpensive and abundant resource of plant proteins can be transformed into a high-tech product for biomedical applications. Of special interest will be to test the bioactivity, biocompatibility and immunogenicity of food proteins in a high-tech application, in which they are not eaten but rather implanted. The use of this economically sound resource is in its infancy, and many parameters for scaffold fabrication from ‘green’ alimentary proteins must be worked out before this approach can find its way into the clinic.

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1

Strength (MPa)

26% 13% 0.1

0%

0.01 % PVOH combined with wheat gluten (a)

100

Elongation at break (%)

26%

10 13% 0%

1

0.1

% PVOH combined with wheat gluten (b)

17.18 Mechanical properties of: (a) tensile stress (MPa); and (b) % elongation-at-break of non-woven fibrous mats derived from wheat gluten as a function of PVOH content: 0% (w/w) PVOH; 13% (w/w) PVOH; 26% (w/w) PVOH. The error bars represent +/– 1 standard deviation (Woerdeman et al., 2007b).

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Cellular responses to electrospun scaffolds: Does fiber diameter matter?

One of the main advantages of electrospinning is the possibility of generating fibrous mats and scaffolds with fiber diameters that can vary by more than 2 orders of magnitude, ranging from below 100 nanometers (Li et al., 2005; Venugopal et al., 2007) to tens of micrometers (Viswanathan et al., 2006) depending on the material used and the electrospinning conditions. While it has been clear from the onset that the biochemical properties of the ensuing scaffolds, as those of any substrate, are of utmost importance in providing appropriate instructional cues, the one major question that has been asked ever since electrospinning was introduced as a biomedical platform technology is: why bother about ‘nano’ – in other words, does the size of fiber diameters matter? It is only recently that some ground-breaking work unequivocally demonstrated that cells sense and respond to the physical properties of the substratum in general (for a recent review, see Discher et al., 2005). Specifically, the fiber size/diameter of electrospun scaffolds is an important instructive cue, as reflected by both the cell morphology (Li et al., 2006a,b) and state of differentiation (Yang et al., 2005). Moreover, tissue-specific cell differentiation and morphology of the ensuing tissue constructs can also be modulated by the fiber density, i.e. the porosity of the scaffolds (Powell and Boyce, 2007). Some recent applications try to combine the beneficial effects of nanoarchitecture (instructive cues at the cell/receptor level) with those of a microfibrous structure, which provides the necessary mechanical strength, for example for bone TE (Tuzlakoglu et al., 2005). The question as to whether to build tissues along aligned (anisotropic) or non-aligned (randomly oriented) fibrous scaffolds remains to be resolved for each tissue under consideration. For example, in neural engineering, an aligned fibrous mat can provide excellent guidance for 3-D axonal growth (Yang et al., 2005). Similarly, in muscular TE, fiber alignment enhances myogenesis in vitro by providing the necessary directional cues along with architectural and mechanical support (Riboldi et al., 2007). As seen in this brief survey, yes, physical features of the scaffolds, such as fiber size (and alignment) do matter, but one size does not fit all! Fiber diameter is obviously a critical, tissue-specific, tunable design parameter that can be effectively utilized for optimally mimicking and emulating the properties of native ECM. What size is ‘desirable’ will depend on the specific cell and tissue type and application in question.

17.7

Conclusions and future trends

In this chapter we have focused on electrospinning of natural polymers as a versatile platform technology for generating fibrous scaffolds for TE with © 2008, Woodhead Publishing Limited

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wide-ranging applications, from vascular grafts to neuronal and cardiac tissues, to bone and cartilage. Given the plethora of innate instructive cues residing in natural mammalian ECM proteins, either the holo-proteins or their enzymatic degradation products (Yamada and Kleinman, 1992; Rachfal and Brigstock, 2005; Kubota and Takigawa, 2007; Badylak et al., 2007), the choice of natural over synthetic proteins as basic ingredients of electrospun scaffolds is obvious. Our data suggest that the specific parameters for spinning optimal tissue-specific scaffolds have to be worked out, one by one, based on the intended applications. Judging from what is known, to date, on the specific interactions of natural scaffolds with cells and tissues in vitro and in vivo, there is sufficient evidence in the literature to unequivocally state that ‘size does matter’ but that ‘one size does not fit all’. Future work will have to focus on: (a) improving the economy of electrospinning by providing better collimation of the solution jet to avoid waste of valuable resources (mammalian ECM proteins don’t come cheap); (b) develop methods for electrospinning more complex ECM analogs that reflect the variety and complexity of the constituents of specific tissues and organs; (c) develop complex electrospinning capabilities for generating scaffolds that reflect the varied 3-D physical and spatial complexity and unique requirements of specific tissues, such as fiber alignment and controlled gradients of porosity; and (d) develop new biomimetic materials by blending natural and synthetic polymers, which in addition to the biological cues of natural proteins and polymers will encompass tunable mechanical properties as well as controlled drug release capabilities.

17.8

Sources of further information and advice

The field of electrospinning of natural proteins for TE scaffolds is very much in flux, and, based on the number of publications, it is certainly in its exponential phase. Specifically, new concepts, materials and technologies are emerging, while established ones are constantly refined. Therefore, any recommendation of key text books or monographs is more or less futile, since by the time a book is published, much of the contents might be outdated. The only exceptions to this rule are books that are published electronically and constantly updated. To the best of our knowledge, the only one such relevant publication is the Encyclopedia of Biomaterials and Biomedical Engineering (Informa Healthcare, publishers). Rather, the reader is referred to the growing number of peer-reviewed specialty journals (some of them are listed here alphabetically), which are the prime venues for publishing relevant quality papers: • • •

Acta Biomaterialia (Elsevier) Biomacromolecules (ACS Publications) Biomaterials (Elsevier)

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• • • • •

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Journal of Biobased Materials and Bioenergy (American Scientific Publishers) Journal of Biomaterials Science, Polymer Edition (Brill) Journal of Biomedical Materials Research (Wiley Interscience) Journal of Tissue Engineering and Regenerative Medicine (Wiley Interscience) Tissue Engineering (Mary Ann Liebert Publishers, Inc)

Finally, the most general advice for anyone interested in the topic is to carry out frequent PubMed searches with a wide spectrum of keywords, encompassing the complex fields of electrospinning, (natural) proteins, TE and scaffolds.

17.9

References

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Teo W E, He W and Ramakrishna S (2006), ‘Electrospun scaffold tailored for tissuespecific extracellular matrix’, Biotechnol, 1(9), 918–29. Torres-Giner S, Gimenez E and Lagaron J M (2008), ‘Characterization of the morphology and thermal properties of Zein Prolamine nanostructures obtained by electrospinning’, Food Hydrocolloids, 22, 235–44. Toshima M, Ohtani Y and Ohtani O (2004), ‘Three-dimensional architecture of elastin and collagen fiber networks in the human and rat lung’, Arch Histol Cytol, 67(1), 31– 40. Tuzlakoglu K, Bolgen N, Salgado A J, Gomes M E, Piskin E and Reis R L (2005), ‘Nanoand micro-fiber combined scaffolds: a new architecture for bone tissue engineering’, J Mater Sci Mater Med, 16(12), 1099–104. Venugopal J, Low S, Choon A T and Ramakrishna S (2007), ‘Interaction of cells and nanofiber scaffolds in tissue engineering’, J Biomed Mater Res B Appl Biomater, 84B(1), 34–48. Viswanathan G, Murugesan S, Pushparaj V, Nalamasu O, Ajayan P M and Linhardt R J (2006), ‘Preparation of biopolymer fibers by electrospinning from room temperature ionic liquids, Biomacromolecules, 7(2), 415–18. Wall J S (1979), In: Recent Advances in the Biochemistry of Cereals, Laidman D L, WynJones R G, (Eds), London, Academy Press, 275–311. Weisenberger M C, Grulke E A, Jacques D, Rantell T and Andrews R (2003), ‘Enhanced mechanical properties of polyacrylonitrile/multiwall carbon nanotube composite fibers’, J Nanosci Nanotechnol, 3(6), 535–9. Woerdeman D L, Veraverbeke W S, Parnas R S, Johnson D, Delcour J A, Verpoest I and Plummer C J (2004), ‘Designing new materials from wheat protein’, Biomacromolecules, 5(4), 1262–9. Woerdeman D L, Ye P, Shenoy S, Parnas R S, Wnek G E and Trofimova O (2005), ‘Electrospun fibers from wheat protein: investigation of the interplay between molecular structure and the fluid dynamics of the electrospinning process’, Biomacromolecules, 6(2), 707–12. Woerdeman D L, Shenoy S and Breger D (2007a), ‘Effects of hydroxyl groups vs. physical entanglements on the electrospinning behavior of wheat protein’, J Biobased Mater & Bioenergy, 1(1), 31–6. Woerdeman D L, Shenoy S and Breger D (2007b), ‘Role of chain entanglements in the electrospinning of wheat protein-poly(vinyl alcohol) blends’, The Journal of Adhesion, 83(8), 785–98. Wool R P and Sun X S (2005), ‘Biobased Polymers and Composites’, Elsevier Academic Press, Boston, 9, 292–326. Yamada Y and Kleinman H K (1992), ‘Functional domains of cell adhesion molecules’, Curr Opin Cell Biol, 4(5), 819–23. Yamauchi F, Yamagishi T and Iwabuchi S (1991), ‘Molecular understanding of heat induced phenomena of soybean proteins’, Food Reviews International, 7, 283–322. Yang F, Murugan R, Wang S and Ramakrishna S (2005), ‘Electrospinning of nano/micro scale poly(L-lactic acid) aligned fibers and their potential in neural tissue engineering’, Biomaterials, 26(15), 2603–10. Yao C, Li X and Song T (2007a), ‘Electrospinning and Crosslinking of Zein Nanofiber Mats’, J Appl Polymer Sci, 103(1), 380–5. Yao C, Li X and Song T (2007b), ‘Fabrication of zein/hyaluronic acid fibrous membranes by electrospinning’, J Biomater Sci Polym Ed, 18(6), 731–42.

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Ye Q, Zund G, Benedikt P, Jockenhoevel S, Hoerstrup S P, Sakyama S, Hubbell J A and Turina M (2000), ‘Fibrin gel as a three dimensional matrix in cardiovascular tissue engineering’, Eur J Cardiothorac Surg, 17(5), 7–91. Zhang X, Min B G and Kumar S (2003), ‘Solution spinning and characterization of poly(vinyl alcohol)/soybean protein blend fibers’, J Appl Polymer Sci, 90(3), 716–21. Zhang Y Z, Venugopal J, Huang Z M, Lim C T and Ramakrishna S (2005), ‘Characterization of the surface biocompatibility of the electrospun PCL-collagen nanofibers using fibroblasts’, Biomacromolecules, 6(5), 2583–9. Zhang Y, Ghasemzadeh S, Kotliar A M, Kumar S, Presnell S and Williams L D (1999), ‘Fibers from Soybean Protein and Poly(vinyl alcohol)’, J Appl Polymer Sci, 71(1), 11–9.

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Part IV Naturally-derived hydrogels: Fundamentals, challenges and applications in tissue engineering and regenerative medicine

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18 Hydrogels from polysaccharide-based materials: Fundamentals and applications in regenerative medicine J. T. O L I V E I R A and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

18.1

Introduction: Definitions and properties of hydrogels

A hydrogel is a network of polymer chains with great water absorbance ability. This implies that once put in an aqueous medium it is able to absorb water and swell, retaining the volume of absorbed water entrapped in the polymeric network. Hydrogels may be formed by a simple reaction between one or more monomers or by association bonds such as hydrogen bonds, van der Waals interactions, among others (Ratner et al., 1996). The chemical composition, crosslinking density, and hydrophobicity can make them vary in consistency from viscous fluids to rigid solids (Varghese and Elisseeff, 2006). Hydrogels may be classified according to different parameters. Peppas (Ratner et al., 1996) determined three major groups in which they can be included, being ranked according to: (1) their method of preparation; (2) ionic charge; or (3) physical structure: 1. Based on the preparation method, they can be homopolymer hydrogels, if constituted by one type of hydrophilic monomer unit; copolymer hydrogels, if consisting of two comonomer units, one of which must be hydrophilic; multipolymer hydrogels, if produced from three or more comonomers reacting together; and interpenetrating polymeric hydrogels, if an initially formed network reacts to form a second intermeshing network structure. 2. Based on the ionic charges, Peppas (Ratner et al., 1996) relates that neutral, anionic, cationic, and ampholytic hydrogels can be formed. Such notation refers to their overall charge, namely no charged groups in neutral hydrogels, negatively charged groups in anionic hydrogels, positively charged groups in cationic hydrogels, and both negatively and positively charged groups in ampholytic hydrogels, which render them with dual behavior. These last three types (anionic, cationic, ampholytic) are also described as ionic hydrogels (or polyelectrolytes). 485 © 2008, Woodhead Publishing Limited

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3. Based on the physical structure, hydrogels can be classified as amorphous hydrogels, if they are non-crystalline containing randomly arranged macromolecular chains; semicrystalline hydrogels, if they include a mixture of amorphous and crystalline phases possessing dense regions of ordered macromolecular chains; hydrogen-bonded structures, if the hydrogel network is based on electrostatic interactions. The strength of the hydrogen bonding is weaker than covalent bonding but stronger than van der Waals interactions. For a hydrogel to maintain its three-dimensional structure, the polymer chains are usually crosslinked chemically or physically. Chemically crosslinked or permanent hydrogels possess their polymeric chains connected by covalent bonds which present a difficulty if one intends to change the shape of these networks afterwards. Physical or reversible hydrogels are linked through non-covalent bonds, such as van der Waals interactions, ionic interactions, hydrogen bonding, or hydrophobic interactions (Figure 18.1). (Hoffman, 2001, Brandl et al., 2007, Varghese and Elisseeff, 2006). Hydrogels can present many different physical forms including solid molded forms, such as soft contact lenses, pressed powder matrices, as in pills or capsules, or different types of microparticles and coatings, or liquids, that form gels upon cooling or heating (Hoffman, 2001). Therefore, their potential application in a diverse range of situations is quite high.

Chemical hydrogel

Crosslinking agent

18.1 Representation scheme of physical and chemical hydrogels formation.

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Different types of natural and synthetic materials are currently studied and used for applications in the tissue engineering and regenerative medicine field (Oliveira et al., 2007b, Brandl et al., 2007). Depending on their role, they can be processed in different ways that range from melt-based technologies, such as fibre extrusion coupled with fibre bonding to generate 3D scaffolds, to solvent based technologies, like hyaluronic acid sponge processing (Oliveira et al., 2007b; Kubo and Kuroyanagi, 2003; Gomes et al., 2002; Moroni et al., 2006). These are expected to create a particular advantage, be it a better fitting ability, or a higher porosity for cells to penetrate, and generate a functional engineered tissue. The processing of a material in a hydrogel is also quite appealing and interesting. The properties of hydrogels make them attractive agents for a diverse range of biomedical applications, such as drug delivery vehicles or cell encapsulating systems (Jen et al., 1996; Coviello et al., 2006). They can be engineered for selective permeability, and mechanical or chemical stability for example. Chemical signals (e.g. pH and ionic factors) and physical stimuli (e.g. temperature or electrical potential) can change the molecular interactions between polymer chains. Such interactions can alter material properties like solubility, swelling behaviour, redox state and crystalline/amorphous transition (Prabaharan and Mano, 2006). The associated biocompatibility, often a result of their hydrophilicity, is also a commonly referred advantage for their use in biomedical and pharmaceutical applications (Jen et al., 1996; Park et al., 1993; Ratner et al., 1996). To this add the good transport of nutrient to cells and products from cells that is normally assured. The ability some possess to be modified with specific cell adhesion ligands or be used as injectable systems further reinforces their potential. Associated disadvantages are usually related to difficult handling and mechanical weakness (Hoffman, 2001).

18.2

Applications of hydrogels produced from different polysaccharides in tissue engineering and regenerative medicine

Polysaccharides are widely distributed in nature in various forms. They include from cellulose, the most predominant polymeric material in nature, to gellan gum, produced by a bacterium strain (Malafaya et al., 2007). Polysaccharides are mainly regarded as sources of energy, starch being an example, but their diverse molecular and physicochemical properties have granted them a place as objects of study in different fields of research. Most polysaccharides can form hydrogels due to their intrinsic properties. Their application in the biomedical field carries great potential due to their chemical behaviour, as well as interesting structural similarities with biological molecules. Research on polysaccharide based materials has been increasing

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in many areas and they are expected to play a key role in future biomedical technology. In the following part of the chapter, some insights on basic aspects of the nature and structure of several polysaccharides will be given. Moreover, potential applications in the Tissue Engineering and Regenerative Medicine field will be described. The gathered information is expected to provide a wider knowledge on polysaccharide based matrices used in the biomedical area, and also increase the interest on this subject enabling other future applications.

18.3

Agarose

Agarose is a linear polymer extracted from marine red algae constituted by (1→3)-β-D-galactopyranose-(1→4)-3,6-anhydro-α-L-galactopyranose units. It is one of the two components that form agar, being the fraction with the greatest gelling capability (Renn, 1984). Agarose gel formation appears to occur by the cooling of an agarose homogeneous solution below the coilhelix transition temperature. The gel is formed when an infinite 3D network of agarose fibers, formed by helices, is developed (Normand et al., 2000). This mechanism was previously suggested by different authors, who suggested that the occurrence of double helices during cooling were responsible for the aggregation that produces the 3D hydrogel network due to hydrogen bonding and hydrophobic interactions (Anderson et al., 1969; Arnott et al., 1974). Nevertheless, other works have suggested that single chain formation would enable gel formation (Foord and Atkins, 1989; Guenet et al., 1993), or the assembly of ternary complexes consisting of agarose-water-cosolvent would lead to the same structure (Ramzi et al., 1996). The melting of agarose gels can occur at higher temperatures, normally around 85°C (Normand et al., 2000). Agarose can be processed without the use of harsh reagents in a relatively clean and simple process, and is non-toxic (Luo and Shoichet, 2004). It is a neutral polysaccharide which may bring some advantages in terms of noninterference with other materials or living tissues with which it may have contact (Renn, 1984). Its gelling kinetics allow homogeneous distribution patterns in cell/drug encapsulation technologies (Haisch et al., 2002). Some drawbacks include its lack of shapable stability, and poorly investigated biodegradability. Some concerns also relate with infectious security and lack of biocompatibility (Haisch et al., 2002). Agarose is commonly used in biochemistry and molecular biology techniques, such as gel electrophoresis or the study of chemotaxis (Nelson et al., 1975). In the regenerative medicine field, agarose hydrogels have been mainly used in the engineering of cartilaginous tissues. A work by Benya et al. used

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rabbit articular chondrocytes cultured in agarose gels where they were shown to reexpress the differentiated phenotype lost due to dedifferentiation in monolayer culture (Benya and Shaffer, 1982). In another study, Mauck et al. (2006) evaluated the chondrogenic differentiation and functional maturation of bovine mesenchymal stem cells and bovine articular chondrocytes in long-term agarose culture. Even though chondrogenesis occurred in mesenchymal stem cells, the amount of the forming matrix and measures of its mechanical properties were found to be lower than those produced by the chondrocytes. Studies on the use of agarose hydrogels for neural regeneration have also been conducted (Stokols et al., 2006). Balgude et al. (2001) revealed that they were able to organize, support, and direct neurite extension from different types of neural cells. These occurrences were also found to occur in a gel concentration-dependent manner (Bellamkonda et al., 1995). In another work, an agarose hydrogel modified with a cysteine compound allowed its patterning with biochemical cues; in this case, the adhesive fibronectin peptide fragment, glycine-arginine-glycine-aspartic acid-serine (GRGDS) was immobilized. In vitro guidance effects of GRGDS oligopeptide-modified channels on the 3D cell migration and neurite outgrowth were observed (Luo and Shoichet, 2004). Another potential application for this material was tested in the creation of a bioartificial pancreas. Iwata et al. showed that the encapsulation of pancreatic islets into agarose microbeads effectively prolonged their functioning in a diabetic mouse and even in a NOD mouse, an animal model of human type I diabetes, without using any immunosuppressive drug (Iwata, 1992).

18.4

Alginate

Alginate is a natural anionic polysaccharide found in seaweed, which is composed of (1-4)-linked β-D-mannuronic acid and α-L-guluronic acid units. The alginate molecule is constituted by regions of sequential mannuronic acid units, guluronic acid units, or by a combination of both. The nature of the alginate dictates the amount and distribution profile of these units (Rowley et al., 1999; Varghese and Elisseeff, 2006). This material can gel in the presence of a small quantity of divalent cations (like Ca2+ or Ba2+) that interact with the carboxylic groups present in the alginate backbone. These groups are present in the guluronic acid residues that, when in contact with those ions, form an ‘egg-box’-shaped structure giving rise to the hydrogel (Varghese and Elisseeff, 2006). The formed hydrogel can be easily disrupted using a chelating agent (e.g. sodium citrate) which captures the cations that maintain the structural integrity of the network (van Osch et al., 1998). Alginate has a well characterized structure which allows for a range of

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comparative studies to be performed, and allows chemical modification through the carboxylic groups in its guluronic acid residues (Rowley et al., 1999). This last feature enabled its lack of cell recognition signals to be overcome by chemically binding RGD peptides, which are extremely important regarding cell-material interactions (Rowley et al., 1999). Alginate hydrogels have also been described to possess low cytotoxicty (Drury and Mooney, 2003). On the other hand, the manufacturing process for the extraction of this polysaccharide from contaminated seaweed leads to the presence of mitogenic, cytotoxic and apoptosis inducing impurities in the final processed material. Even though such molecules can be removed by further purification steps, it is a time and money consuming process. Also, variations in the mannuronic and guluronic acid composition in each individual sample confer variability to the processed samples (Leinfelder et al., 2003). This is an important issue since it has been shown that alginates with a high guluronic acid content develop a great inflammatory response (De Vos et al., 2002). Other drawbacks are common to most polysaccharide hydrogels and are related to weak mechanical properties and uncontrollable degradation profile (Varghese and Elisseeff, 2006) In terms of applications, alginate is commonly used as a mold-making material in dentistry, prosthetics, textiles, and in the food industry, for thickening soups and jellies (Sharon-Buller and Sela, 2007). Concerning tissue engineering, alginate is well known in cartilage regeneration approaches. It has been used to encapsulate human articular chondrocytes and cultured in the presence of recombinant human BMP-2, which is revealed to have positive effects on collagen type II expression (Grunder et al., 2004). In another study, human adipose tissue-derived stromal cells were encapsulated in alginate hydrogels and were shown to produce characteristic cartilage matrix molecules in both in vitro and in vivo models (Erickson et al., 2002). Other works focused on tissues such as bone, skin and heart, for example. Alsberg et al. (2001) cultured primary rat calvarial osteoblasts with alginate modified hydrogels observing significant increases in bone formation in vivo. Wang et al. showed that alginate hydrogels were able to support proliferation and differentiation of rat bone marrow cells along the osteoblastic lineage (Wang et al., 2003). Skin regeneration has been experimented using sponges composed of gelatine and alginate (Choi et al., 1999); and alginate macrobeads containing pancreatic islets had a positive effect on diabetes reversal (Trivedi et al., 2001). Liver tissue engineering was also attempted by combining hepatocytes with alginate scaffolds (Dvir-Ginzberg et al., 2003). As a final example, alginate scaffolds in which cardiomyocytes were seeded showed spontaneous contraction in some of the aggregates as well as the ability to allow coculturing of both cardiomyocyes and cardiofibroblasts (Dar, 2002). This

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opens interesting alternatives concerning the use of this material for cardiac tissue renewal.

18.5

Carrageenan

Carrageenans are a family of linear, water soluble, sulfated, anionic polysaccharides extracted from marine red algae. They are large and highly flexible molecules that form a variety of different gels at room temperature due to the assembly of helical structures (Rees, 1972; Mangione et al., 2003). Different types of carrageenans can be distinguished based on their primary structure. Kappa-carrageenan or κ-carrageenan, is composed of alternating α-(1-3)-D-galactose-4-sulfate and β-(1-4)-3,6-anhydro-D-galactose; iotacarrageenan, or ι-carrageenan, differs from this by the presence of an additional sulphate group at C2 of the 1,4 linked galactose unit, while lambda-carrageenan, or λ-carrageenan, presents a third sulfate group at C6 of the 1,4 linked galactose unit. They form gels with different features: κ-carrageenan gives rise to strong, rigid gels; ι-carrageenan forms soft gels, and λ-carrageenan forms gels when mixed with proteins rather than water (Mangione et al., 2003). In aqueous environment and in the presence of cations, both κcarrageenan and ι-carrageenan form thermoreversible gels on cooling, whereas gelation of λ-carrageenan has not been observed (Mangione et al., 2003). Pioneering work by Rees et al. gave great insights on the conformational and structural properties of κ-carrageenan (Rees, 1972). The gelation of κcarrageenan involves, upon cooling, a coil-double helix transition followed by the aggregation of helice domains giving rise to large molecular aggregates that constitute the κ-carrageenan network (Evageliou et al., 1998; Mangione et al., 2003). The helical structures are further stabilized with hydrogen bonds perpendicular to the helix axis (Rees, 1972). The melting of these gels can be attained by increasing temperature. An initial change in helical conformation affecting the junction zones causes network disruption followed by melting of aggregates and consequent conformational helical change, when the temperature is further increased. Several factors influence gel formation, such as polysaccharide chemical structure, nature of co- and counterions, polymer concentration, and temperature (Meunier et al., 2001; Mangione et al., 2003). One of the main characteristics of carrageenans is their thixotropy – which means they thin under shear stress and recover their viscosity once the stress is removed. This enables them to be easily injected but adopt a solid shape in the mold afterwards. It is also a material that can be processed using nonharsh reagents under mild conditions. A major drawback is the carrageenan induced inflammation that may hinder its use in the biomedical area. Carrageenan is commonly used to induce chronic inflammatory arthritis in

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various animal models (Hansra et al., 2000; Aloe et al., 1992; Erel et al., 2004). Furthermore, a recent publication indicates that carrageenan induces inflammation in human intestinal epithelial cells in tissue culture through a Bcl10-mediated pathway that leads to activation of NF-κB and IL-8 (Borthakur et al., 2007). Carrageenan is mostly used for pharmaceutical and environmental applications, as well as drug delivery approaches to some extent (Bonferoni et al., 2004; Sjoberg et al., 1999; Leung et al., 1995; Guo et al., 1998). Its use in tissue engineering approaches is not widespread in the literature. Even so, it has been combined with collagen and hydroxyapatite in order to create an injectable substitute to be used in bone tissue engineering approaches (Gan and Feng, 2006).

18.6

Cellulose

Cellulose is the most widespread polymeric material in nature. The most common is a fibrous, tough, water-insoluble material that can be found in the cell walls of plants, mainly in stalks, stems or trunks. In addition, cellulose from bacterial origin structurally similar to the cellulose produced by plants can also be found in microorganisms (Hestrin, 1962; Colvin, 1980; Wong et al., 1990). The gram-negative bacterium Acetobacter xylinum has been used for this purpose (Hestrin, 1962; Colvin, 1980). It produces highly pure cellulose that resembles in crystallinity and average microfibrillar width that from many plants and algae, while presenting extensive polymerization ability (Ross et al., 1991). It constitutes a good alternative to plant and algae cellulose being also applied in different fields (Klemm et al., 2001; Legeza et al., 2004; Svensson et al., 2005). Cellulose is composed of β-D-glucan units linked by (1→4) glycosidic bonds that are formed by a simple polymerization of glucose residues from a substrate such as UDP-glucose (O’Sullivan, 1997; Delmer and Amor, 1995). The stereochemistry conferred by the glycosidic linkage creates a linear extended glucan chain that enables a precise and ordered interaction between different chains. This material exists as a combination of various chains strongly linked by hydrogen bonding, named microfibrils, instead of a single chain, which contributes to its rigid structure (Delmer and Amor, 1995). The hydroxyl groups that hold the cellulose chains together account for the high degree of crystallinity, low solubility, and poor degradation in vivo. Even though bacterial cellulose contains approximately 90% water as prepared, this water is easily squeezed out and the recovery in the swelling property is complicated by the hydrogen bonds (Nakayama, 2004). Cellulose presents different polymorphs, cellulose I, or native cellulose, being the more common in nature (O’Sullivan, 1997). Cellulose possesses high strength in the wet state, and is biocompatible.

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Bacterial cellulose exhibits a rapid growth and the ability to be maintained under controlled conditions, which is another advantage (Ross et al., 1991). On the other hand, the inherent low water solubility may compromise its degradation profiles both in vitro and in vivo (Ross et al., 1991). In addition, it is not biodegradable due to lack of digestive enzymes in the human organism (Miyamoto et al., 1989, Hayashi, 1994). Cellulose has its major applications in the paper and textile industries (Princi et al., 2006). Concerning the biomedical field, frequently cellulose derivatives, such as methylcellulose, hydroxypropylcellulose and carboxymethylcellulose, are used as starting materials due to the new functionalities gained upon these modifications (Gutowska et al., 2001). Applications of cellulose based materials have been reported for different tissues, such as cartilage and skin. In a study by Svensson et al. (2005), bacterial cellulose supported bovine chondrocytes’ proliferation to some extent, while providing significant advantages in terms of mechanical properties. In following studies, human chondrocytes proliferation was supported by cellulose based scaffolds. A bacterial cellulose hydrogel wound dressing was used in the reparative processes of deep dermal burns by being impregnated with wound therapeutic agents. Results showed that some healing of the burn wounds occurred (Legeza et al., 2004). Moreover, a preliminary study was performed to create artificial blood vessels for microsurgery applications using bacterial cellulose based systems (Klemm et al., 2001).

18.7

Chitin/chitosan

Chitin occurs in a wide variety of species, arthropod shells being the most accessible sources. Chitin is structurally similar to cellulose, and is a linear, high molecular weight crystalline polysaccharide consisting of β-(1→ 4) linked D-glucose, the basic repeating unit being N-acetyl-D-glucosamine. Three crystalline forms are known (α, β and γ) from which α-chitin is the most common type. α-chitin possesses its molecules aligned in an antiparallel fashion that grants them strong molecular hydrogen bonding and structural stability. The β-type, present in squid pens for example, has a parallel molecular arrangement that confers weaker intermolecular forces. Such a feature may be interesting for some applications. γ-chitin is the less common and is considered to be a mixture of the other two types, presenting mixed antiparallel and parallel molecular arrangements (Kurita, 2001). Chitosan is often more popular than chitin in some biomedical applications due to its intrinsic features such as hydrophilicity, or ready solubility in dilute acids which renders it more accessible for chemical reactions (Khor and Lim, 2003). Chitosan is a semi-crystalline cationic polysaccharide that results from the partial alkaline deacetylation of chitin. It is constituted by β-

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1,4-linked 2-amino-2-deoxy-D-glucose and is normally insoluble in aqueous solutions above pH 7 (Freier et al., 2005a). However, in dilute acids (pH 6), the free amine groups are protonated and the molecule becomes soluble. This pH-dependent solubility provides a convenient mechanism for processing under mild conditions (Madihally and Matthew, 1999). These polymers can be easily produced due to the high annual production and great accessibility of chitin. They have been described as non-toxic, biodegradable and have antibacterial activity (Sashiwa and Aiba, 2004). They are also biocompatible, while possessing structural similarities to glycosaminoglycans, which are structural components of the cartilage extracellular matrix (Mahmoudifar and Doran, 2005; Correlo et al., 2005; Drury and Mooney, 2003). Associated disadvantages include the intractable bulk structure of chitin and the weak mechanical properties often exhibited by chitosan. Moreover, these polymers have been described to possess neutrophile recruitment ability, a feature normally linked with acute inflammation (Usami et al., 1994). Even so, this appears not to influence their biocompatibility (VandeVord et al., 2002). Chitin is frequently used in water purification, and as an additive to thicken and stabilize foods and pharmaceuticals (Knorr, 1982). Chitosan serves different applications and its use ranges from the food industry to the biomedical and pharmaceutical fields (Devlieghere et al., 2004; Silva et al., 2003, Singh, 2000). Their use has been increasing over the last few years in the biomedical field, and studies on bone, cartilage and neural regeneration can be found. Hydroxyapatite-chitin porous matrices were cultured with rabbit mesenchymal stem cells and induced into osteoblasts in vitro using dexamethasone. These osteoblasts were cultured on the matrices and implanted into bone defects of the rabbit femur. Histology of explants showed bone regeneration with biodegradation of the hydroxyapatite-chitin scaffolds (Ge et al., 2004). In a study by Seol et al. (2004), rat calvarial osteoblasts were grown in porous chitosan sponges and the results showed that bone formation occurred. In another study, Tuzlakoglu et al. (2004) produced 3D fiber mesh chitosan scaffolds that were seeded with osteoblasts and, upon culturing, presented good morphology and no inhibition of cell proliferation. Concerning cartilage tissue engineering, Chenite et al. (2000) combined chitosan with polyolphosphate salts to produce an injectable system. This was successfully used to deliver biologically active growth factors in vivo, and also to encapsulate chondrocytes for tissue engineering applications. Focusing on osteochondral repair strategies, Oliveira et al. (2006) used hydroxyapatite/chitosan bilayered scaffolds to support the growth and differentiation of goat marrow stromal cells into osteoblasts and chondrocytes. Also on the same subject, Malafaya et al. (2005) developed chitosan particle agglomerated scaffolds and seeded them with human adipose derived mesenchymal stem cells. These were cultured with conditioned mediums to enable their differentiation towards

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the chondrogenic and osteogenic lineages. Preliminary studies indicate the cells were adopting both osteogenic and chondrogenic morphologies in each predestined part of the 3D particle agglomerated scaffolds. Neural regeneration approaches have also been attempted using chitin and chitosan hydrogel tubes. These were shown to support adhesion and differentiation of primary chick dorsal root ganglion neurons in vitro (Freier et al., 2005b). As a final example, several applications for wound dressing have been suggested. An asymmetric chitosan membrane has been studied in this sense and histological examination confirmed that epithelialization rate was increased and the deposition of collagen in the dermis was well organized (Mi et al., 2001).

18.8

Chondroitin sulfate

Chondroitin sulfate is a sulfated glycosaminoglycan (GAG) composed of repeating disaccharide units of D-glucuronic acid and N-acetylgalactosamine (Malafaya et al., 2007; Wang et al., 2007). These sugars can exist in wide extensions in a chondroitin chain and be sulfated in different positions. This molecule is an important structural component of the extracellular matrix of cartilaginous tissues, and by binding core protein gives rise to aggrecan, the most important proteoglycan in cartilage. The tightly packed and highly charged sulfate groups of chondroitin sulfate generate electrostatic repulsion that provides much of the resistance of cartilage to compression (Roughley et al., 2002). This cooperates in the functioning of aggrecan as a shock absorbing molecule (Malafaya et al., 2007). Sources for chondroitin sulfate include extracts of cow trachea and pig ear and nose cartilage, although shark cartilage may also be used (Mauck et al., 2000). Chondroitin sulphate is quite water soluble, which limits its use alone in the solid state for biomedical applications, being frequently combined with other polymers (Li, 2004; Lee et al., 2000). In fact, its anionic nature enables efficient interaction with cationic molecules to form interesting structures (Sechriest, 2000). Its overall ability to function as a cell interacting molecule has spread its use in a variety of biomedical applications. Chondroitin sulfate is currently used as an ingredient in dietary supplements, with the ultimate goal of relieving some of the pain and disability of patients with musculoskeletal pathologies, namely osteoarthritis. Even so, none of its beneficial effects when compared to control groups has been proved, according to Clegg et al. (2006). Due to its nature, chondroitin sulfate has been mostly used in the development of supports for cartilage tissue engineering applications. A work performed by Sechriest et al. (2000) combined chondroitin sulfate and chitosan, in order to develop a new biomaterial for chondrogenesis support. Bovine articular chondrocytes seeded onto chondroitin sulfate-chitosan membranes were shown to form focal adhesions and maintain characteristics of

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differentiated chondrocytes, including collagen type II and proteoglycan production. In another study, chondroitin sulfate was covalently attached to a type I collagen porous scaffold in order to observe the changes on the metabolic activity of seeded chondrocytes. It was shown that chondroitin sulfate positively influenced cell proliferation and the total amount of synthesized proteoglycans (van Susante et al., 2001). A study in the bone tissue engineering area used chondroitin sulfatechitosan sponge seeded with rat calvarial cells and cultured for different time periods. Results showed the seeded osteoblasts were well attached and proliferated within the chondroitin sulfate-chitosan sponges (Park et al., 2000). Blood vessel formation was also enhanced in a mouse model by implanting crosslinked glycosaminoglycan-based hydrogels loaded with basic fibroblast growth factor (Cai et al., 2005). As a final example, kidney tissue engineering was approached using chondroitin sulfate based matrices. Polyvinyl alcohol-chondroitin sulfate hydrogels were used to culture baby hamster kidney cell line cells, promoting their effective growth on the surface (Lee et al., 2005).

18.9

Dextran

Dextran is a complex, branched polysaccharide made of many glucose molecules joined into chains of varying lengths. The straight chain consists of α (1→6) glycosidic linkages between glucose molecules, while branches begin from α (1→3) linkages (and in some cases, α1→2 and α1→4 linkages as well). Some dextrans possess almost all the linkages in α (1→6) form, although others only share that feature in 50%. Dextran is synthesized from sucrose by certain lactic-acid bacteria, the best-known being Leuconostoc mesenteroides (Misaki et al., 1980; Jeanes et al., 1948). These dextrans are high molecular weight α-glucans, varying from slightly to highly branched, in which 1→6 linkages are predominant (Cerning, 1990). Nevertheless, other strains and bacterium exist, that also form dextrans from sucrose cleavage (Horisberger, 1969; Lewick et al., 1971). They give rise to dextrans with different extents of branching, their long chain being responsible for the physico-chemical properties. Dextran’s hyperbranched nature confers various options in terms of functionality (Ioan et al., 2000). Dextran hydrogels may be formed by chemically crosslinking the polymer chains using reagents like diisocyanates, and epichlorohydrin, or by derivatization with polymerizable methacrylate groups, for example. Also, physical crosslinking can be reached by ionic or hydrophobic interactions (Stenekes et al., 2001). Dextran is essentially non-toxic (Stenekes et al., 2001). Some of the associated disadvantages relate to the wide variability of the extracted material and the reagents sometimes involved in the hydrogels’ processing that may exert some negative effects on the contacting biological material (Stenekes et al., 2001). © 2008, Woodhead Publishing Limited

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Dextran is commonly used as an antithrombotic and in solution for other medical purposes (Zhang and Wieslander, 1993). In the biomedical field, dextran has been frequently used as a drug delivery system (Cadée, 2001; Malafaya et al., 2002). Even though its use in tissue engineering and regenerative medicine approaches is quite scarce, some studies have been performed commonly using modified dextran (De Groot et al., 2001). Grafting of cell adhesion peptides, like RGD, on low protein-binding dextran monolayer surfaces have stimulated endothelial cell, fibroblast, and smooth muscle cell attachment and spreading in vitro, while also promoting cell type-dependent interactions (Massia, 2001). Such findings may enable the use of these polysaccharide materials for the processing of other structures to be employed in a tissue engineering approach. Research on dextran scaffolds used foams of polylactide-grafted dextran copolymers and seeded them with mouse 3T3 fibroblasts. Results showed the dextran blends had good cell affinity and moderate mechanical strength which might prospect their use as scaffolds for tissue engineering (Cai et al., 2003). In another study, hydrogels made of dextran and hyaluronan showed good cytocompatibility in vitro using vascular smooth muscle cells (Trudel and Massia, 2002).

18.10 Gellan Gellan gum is a linear anionic polysaccharide composed of tetrasaccharide repeating units of 1,3-β-D-glucose, 1,4-β-D-glucuronic acid, 1,4-β-D-glucose, 1,4-α-L-rhamnose, containing one carboxyl side group, and was initially described by Moorhouse et al. (Jansson et al., 1983; Moorhouse, 1981). Two gellan gum forms exist, acetylated – which is the initial product produced by Sphingomonas paucimobilis, and deacetylated – the processed and most common form. They form thermoreversible gels with differences in mechanical properties from soft and elastic for the acetylated form, to hard and brittle for the fully deacetylated polysaccharide (Kang et al., 1982; Grasdalen and Smidsrod, 1987). Gellan gum has an ionotropic gelation, similar to other polysaccharides, like alginate or carrageenan. At high temperatures gellan gum is in the coil form, but upon temperature decrease transits to a double-helix. These helices self-assemble to form oriented bundles, called junction zones. Afterwards, untwined regions of polysaccharide chains link the junction zones leading to the formation of a gel (Quinn et al., 1993). The gelation of gellan gum solutions is strongly influenced by the chemical nature and quantity of cations present in the solution. Divalent cations promote the gelation much more strongly than monovalent cations. In monovalent cations, the gelation is mainly a result of the screening of the electrostatic repulsion between the ionized carboxylate groups on the gellan gum chains. In the case of divalent cations, the gelation and aggregation of gellan occurs via a chemical bonding

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between divalent cations and two carboxylate groups belonging to glucuronic acid molecules in the gellan chains, in addition to the screening effect (Singh and Kim, 2005). Gellan gum can be easily processed into transparent gels that are resistant to heat and acid stress without the use of harsh reagents (Ogawa, 1999). It is not cytotoxic and can be used as an injectable system (Oliveira et al. 2006b). It has been previously used in vivo in human patients as an ocular drug delivery vehicle (Dickstein et al., 2001; Shedden et al., 2001). One major drawback is its weak mechanical properties. Gellan gum gels are commonly used in the food industry as thickening agents or stabilizers (Mao et al., 2000). In the biomedical field, most applications are suggested for drug delivery approaches (Coviello et al., 1998; Kubo et al., 2003). One example is a work by Sanzgiri et al. (1993) which evaluated the efficacy of a methylprednisolone ester of gellan gum for applications in ocular sustained release devices. Few works can be found about the use of gellan gum in tissue engineering approaches. Even so, it has been proposed as a cell support for cartilage regeneration approaches. In a work by Oliveira et al. (2006a), human nasal chondrocytes have been encapsulated in gellan gum hydrogels and cultured for periods of eight weeks, showing that these systems are suitable for supporting chondrocyte development. Later studies compared the suitability of in vivo injectable gellan gum hydrogels to encapsulate and support the culturing of human articular chondrocytes and human bone marrow cells for cartilage tissue engineering applications. Results were in accordance with those previously obtained for human nasal chondrocytes. Moreover, human articular chondrocytes were shown to actively synthesize extracellular matrix components as determined by realtime PCR and histological analysis (Oliveira, 2006b). A work by Ciardelli et al. (2005) blended polycaprolactone with different polysaccharides, including gellan gum. The main conclusions refer that starch and gellan are suitable materials for the production of tissue engineering scaffolds made from PCL based blends. Another work performed by Suri et al. (2006), referred at biopolymers based on gellan gum and hyaluronic acid being tested as in situ gelling short-term vitreous substitutes. The results obtained suggested that these combined systems are suitable as in situ gels, having biophysical properties similar to that of the vitreous gels.

18.11 Hyaluronic acid Hyaluronan is a non-sulfated glycosaminoglycan distributed widely throughout connective, epithelial, and neural tissues. Hyaluronic acid is a linear, negatively charged polysaccharide, constituted by a mixture of two sugars, glucuronic acid and N-acetyl glucosamine. It is linked together via alternating β-1,4 and β-1,3 glycosidic bonds constituting large molecules. Bulky groups on each

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sugar molecule are in sterically favored positions while the smaller hydrogens assume the less favorable axial positions, which confers them energetic stability (Haxaire et al., 2000). Hyaluronic acid may derive from various sources, being the most prevalent rooster combs and Streptococcus bacterium, by recombinant production. They possess different qualities such as variations in rheological properties (Price, 2007). Hyaluronic acid is highly viscous in solution, but its shear dependent viscosity degree allows it to be injected through a small gauge needle (Mallapragada, 2006). This molecule is one of the chief components of the extracellular matrix, contributing significantly to cell proliferation and migration. Hyaluronic acid is soluble in water and is not resistant to enzymatic degradation (Barbucci et al., 2000). The modification of its chemical structure has granted specific attention to this material, the more known being the formation of esters of hyaluronic acid through the esterification of its carboxylic groups (Della Valle and Romeo, 1987; Cortivo et al., 1991). This change enhances its processability and resistance to a range of conditions, enabling processing into membranes, spheres, and different porous structures, for example (Cortivo et al., 1991). Hyaluronic acid has the main advantage of being an important component of the extracellular matrix. Moreover, it has the possibility of being administered as an injectable system and is biocompatible. As with other systems, weak mechanical properties are associated, and connection with malignant tumor progression and use as a tumor marker as been described (Meyer and Chaffee, 1940; Naor et al., 1997). Although hyaluronic acid has been used in the cosmetic industry, it is most widely applied for tissue engineering and related approaches (Cortivo et al., 1991; Lowe et al., 2001). Bone regeneration has been approached by combining acrylated hyaluronic acid with bone morphogenic protein-2 and human mesenchymal stem cells for rat calvarial defects. Results showed an expression of bone markers, such as osteocalcin, as well as mature bone formation with vascular markers (Kim et al., 2007). These materials are also widely used in the cartilage regeneration field (Solchaga, 1999). Hyaluronic acid benzyl ester derived scaffolds (HYAFF®11) were used to culture human nasoseptal chondrocytes. The cells expressed cartilage-specific collagen type II, indicating that they were able to redifferentiate and retain a chondrocytic phenotype (Aigner, 1998) Human fibroblasts have also been seeded on a scaffold of HYAFF®11 and combined with human keratinocytes in order to obtain a dermal–epidermal composite structure. Results showed that human fibroblasts and keratinocytes could be cultured on these materials and that the pattern of expression of particular dermal–epidermal molecules was similar to that found in normal skin (Zacchi, 1998). In addition, fibroblasts have also been used to synthesize new extracellular matrix on hyaluronic acid scaffolds and assess the influence this had on the proliferation rate and survival of rat hepatocytes during longterm culture and after in vivo transplantation. The secreted extracellular

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matrix appears to improve the biological properties of the hyaluronan-based scaffolds, favouring the survival and morphological integrity of hepatocytes both in vitro and in vivo (Zavan et al., 2005). Hyaluronic acid has been studied for heart regeneration and several studies have been conducted toward the tissue engineering of heart valves (Masters et al., 2004; Masters et al., 2005; Ramamurthi and Vesely, 2005). Boublik et al. (2005) have produced hybrid cardiac grafts based on knitted hyaluronic acid fabric and fibrin. Results evidenced that mechanical stimulation led the grafts to exhibit mechanical properties comparable to those of native neonatal rat hearts. Moreover, in a subcutaneous rat implantation model the constructs exhibited the presence of cardiomyocytes and blood vessel ingrowth after three weeks. Focusing on neural tissue engineering, Tian et al. (2005) developed a hyaluronic acid based hydrogel, with an open porous structure and viscoelastic properties similar to neural tissue, and cultured neural cells on these systems. Glial fibrillary acidic protein-positive cells – reactive astrocytes – were found to infiltrate the hydrogel matrix and further tests indicated the compatibility of this type of support with brain tissue.

18.12 Starch Starch is a natural polymer made of a combination of two polymeric carbohydrates, amylose and amylopectin (Han and Lim, 2004; Funami et al., 2005). In terms of relative weight percentages, the amount of amylopectin is higher than amylose in common types of cereal endosperm starches (Buleon et al., 1998). Amylopectin is the highly branched component of starch, being formed of chains of α-D-glucopyranosyl units linked mainly by (1→4) linkages but with 5-6% of (1→6) bonds at the branch points. The overall composition has hundreds of short (1→4)-α-glucan chains interlinked by (1→6)-α-linkages (Buleon et al., 1998; Reis et al., 2001). Amylose is a linear molecule of (1→4) linked α-D-glucopyranosyl units, even though some molecules are slightly branched by (1→6)-α-linkages. Evidence suggests that the branched linkages are frequently located rather near the reducing terminal end and/or they have multiple branched chains (Takeda et al., 1987). Water soluble starches can be dispersed in water and upon heating, form a paste. When a cooling regime is imposed, the starch paste increases in viscosity giving rise to a hydrogel caused by the physical crosslinking of hydrogen bonds (You, 2003). Starch-based polymers are degradable and biocompatible. They exhibit distinct structural forms and properties that can be tailored by the other component of the starch-based blend. Moreover, they are an abundant and low cost product (Gomes, 2003). Some drawbacks associated with this material include its processability, which often implies the use of a synthetic polymer (Salgado et al., 2002).

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Starch is commonly used as a food thickening agent, being also applied in the manufacturing of adhesives, paper, and textiles (Kraak, 1992). In the regenerative medicine field, most of the work conducted on starch-based polymers has been directed toward bone and cartilage tissue engineering applications. Reis and co-workers have performed extensive research on corn starch-based polymers, using different material processing technologies to generate various structures (Gomes et al., 2001; Elvira et al., 2002; Oliveira et al., 2007a). Salgado et al. (2002) used starch/cellulose acetate scaffolds to conduct a preliminary study on the adhesion and proliferation of human osteoblasts. Cell culturing experiments showed that cells were viable and that there were no signs of cellular death after three weeks of culture. Biochemical assays demonstrated that cells maintained the osteogenic phenotype throughout the experiment and deposition of mineralized extracellular matrix could be detected. In another work (Gomes, 2003), also on bone tissue engineering, the effect of culturing conditions (static and flow perfusion) on the proliferation and osteogenic differentiation of rat bone marrow stromal cells seeded on starch-based scaffolds was investigated. It was concluded that flow perfusion culture enhanced the osteogenic differentiation of marrow stromal cells and improved their distribution in the 3D structures. Some studies have also been performed in the cartilage tissue engineering field. Oliveira et al. (2007a) seeded bovine articular chondrocytes onto SPCL fiber mesh scaffolds under dynamic conditions and evaluated their suitability for supporting chondrocyte development and extracellular matrix synthesis. Active cell proliferation was observed, as well as deposition of specific extracellular matrix components, such as collagen type II. Besides bone and cartilage, other cell types and target tissues have been investigated. Ciardelli et al. (2005) used bioartificial blends of polycaprolactone with starch processed into scaffolds by selective laser sintering. These structures were combined with NIH-3T3 mouse fibroblasts to evaluate the rate and the extent of cell adhesion. Starch has also been tested for the promotion of angiogenesis and blood vessel formation. Santos et al. (2007) examined the interaction of both micro- and macrovascular endothelial cells with starch-based scaffolds. Endothelial cells growing on the SPCL fibers exhibited a typical morphology, while maintaining important functional properties, which reveals good prospects in terms of their use for blood vessel formation.

18.13 Xanthan Xanthan is produced by fermentation of glucose or sucrose using Xanthomonas campestris bacterium and has a pentasaccharide repeating unit structure (Jansson et al., 1975). The main chain consists of (1→4)-β-D-glucose units with a terminal β-D-mannose, β-D-glucuronic acid, and β-D-mannose side

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chain which has β-D-(1→2) and →-D-(1→4) linkages (Chellat, 2000a). Some of the terminal mannose units have pyruvic acid attached and the inner mannose is partially acetylated (Sereno et al., 2007). Xanthan has polyanionic properties which enables it to interact with polycationic molecules (Chellat, 2000a). In the solid state and in the presence of a nominal amount of water, xanthan adopts a double helical conformation. Dilute solutions, above a threshold concentration, have the ability to form gels in which structural integrity is maintained by junction zones composed of oriented bundles of xanthan double helices. These junction zones are supposed to be linked by extended helical chains, resulting in a 3D network. Its sol-gel transition depends on the concentration, molecular weight, and nature of cations present, in a similar fashion to other hydrogels. Xanthan can be processed under mild conditions without using harsh reagents. Nevertheless, the preparation as a raw material often affects its molecular weight which introduces a high degree of variability to the obtained samples (Sereno et al., 2007). Xanthan gum is normally used as a liquid viscosifier in foods, finding also some applications in the oil industry as a thickener for drilling fluids (Bertrand and Turgeon, 2007). In the biomedical area, as with most of the polysaccharides mentioned, it has been used as a drug delivery system (Vendruscolo et al., 2005; Bertram and Bodmeier, 2006; Andreopoulos and Tarantili, 2001; Chellat et al., 2000b; Dumitriu and Dumitriu, 1991). Not so much research on tissue engineering applications of xanthan gum has been conducted. Several other polysaccharides have been suggested for applications in the regenerative medicine field, but studies on these biomaterials are quite scarce. Some examples are galactomannans, polysaccharides consisting of a mannose backbone with galactose side groups (Chaubey and Kapoor, 2001); gum arabic, a complex mixture of saccharides and glycoproteins (Ramakrishnan et al., 2007); and pullulan, a polysaccharide composed of maltotriose units (Al-Assaf et al., 2006). In addition to this, others such as levan, elsinan, scleroglucan, and guar gum are mentioned (Shalaby, 1994; Coviello, 2006).

18.14 Conclusion The use of polysaccharides is increasing in the biomedical field, and more precisely in the tissue engineering and regenerative medicine area. It was shown here that several polysaccharides, some well-known and others not so well known, carry enormous potential to be used in repair treatments for a range of tissues that extend from bone to liver. Such interest in this class of natural materials is a result of specific advantages they grant, like non-harsh processing, variable degrees of hydrophilicity, similar chemical structure to living organism molecules, and biocompatibility, among others. In fact, the variety of existing polysaccharides described here opens a wide range of

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opportunities to create new and multifunctional materials by making use of specific features in each of them. One can combine anionic and cationic polymers generating a new structure, confer a determined functionality by binding bioactive agents of interest, or develop support and is drug delivery systems to enhance the ability of these biomaterials to regenerate a chosen tissue. Although active research is being conducted and is increasing, some problems still exist in some systems, such as the easy solubility in water which makes difficult the formation of a stable hydrogel, or the weak mechanical properties that are a common trend among polysaccharide materials. The use of hydrogels for tissue engineering applications also creates the need to look at the overall scenario from the biological side. The different materials used must allow proper conditions for cell viability and function, and once implanted in a living organism should not elicit a severe inflammatory and immune response from the host. These features are of course closely linked to others such as proper permeability to allow efficient oxygen and nutrient diffusion, as well as removal of toxic by-products (Jen et al., 1996; Lee and Mooney, 2001). Other characteristics are related with particular hydrogel uses: for example, if an encapsulation process is conducted it should enable uniform cell distribution and be formed under mild polymerization conditions; also in cell seeding approaches sites that enable or promote cell adhesion and proliferation should be present; as a final example, if the tissue of choice is highly vascularized the fabricated support should cope with those needs promoting efficient blood vessel network formation (Lee and Mooney, 2001; Patrick, 2004). As a summary, it can be stated that one must always consider the material and processing technology used along with the cells to be employed and the final application of the product. The problem should be addressed globally even though specific aspects must be dealt with along the way, creating in this sense regenerative technologies with increased potential at the end of the process. Continuous research and study until now have allowed surmounting several of these barriers and future work will enable the creation of optimised and close to flawless systems. Even though polysaccharide based materials have a strong tradition in drug delivery approaches, their application in the tissue engineering area is growing and will have an important role in future treatments for the improvement of patients’ life quality and wellbeing.

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19 Alginate hydrogels as matrices for tissue engineering H. P A R K and K.-Y. L E E, Hanyang University, South Korea

19.1

Introduction

The number of patients who suffer the loss or failure of an organ or tissue due to either accident or disease is rapidly growing. Tissue or organ transplantation can treat these patients, but donor availability is very limited. Tissue engineering is one promising and potential strategy to provide the needed organs or tissues (Langer and Vacanti, 1993; Lee and Mooney, 2001). In this strategy, tissues or organs can be potentially engineered through a combination of cells and polymer scaffolds. In brief, tissue-specific cells are isolated from a small biopsy from the patient and cultured in vitro. The cells are subsequently incorporated into a three dimensionally-constructed polymer scaffold. This polymer scaffold mimics many roles of extracellular matrices (ECMs), delivers the cells to the desired site, provides a space for new tissue formation, and controls the structure and function of the engineered tissue (Putnam and Mooney, 1996; Marler et al., 1998). Hydrogels are potential scaffolds for tissue engineering, as they have structural similarity to the macromolecular-based components of the body and are biocompatible. In addition, hydrogels can be transplanted into the body in a minimally invasive manner using a syringe or endoscope, which can reduce both procedural pain and recovery time. A number of synthetic and natural polymers have been used to prepare hydrogels for tissue engineering applications. Alginate is a typical biomaterial that has been extensively investigated and used for many biomedical applications due to its biocompatibility, low toxicity, relatively low cost, and mild gelation condition with divalent cations such as Ca2+. In this chapter, the general features of alginate and its potential applications in tissue engineering will be discussed.

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Properties of alginate

19.2.1 General properties Commercially available alginate is typically extracted from brown algae, including Laminaria hyperborean, Ascophyllum nodosum and Macrocystis pyrifera (Smidsrod and Skjak-Bræek, 1990). Bacterial alginates have also been isolated from Azotobacter vinelandii and several Pseudomonas species (Skjak-Bræek et al., 1986). Until Fischer identified the guluronic acid residue (Fischer and Dorfel, 1955), D-mannuronic acid was thought to be the major component in alginate hydrolysates (Nelson and Cretcher, 1929). Fractional precipitation with manganese and calcium salts demonstrated later that alginates are actually block copolymers, and that the ratio of guluronic to mannuronic acid varies depending on the source (Haug, 1959). Alginate is a linear copolymer containing blocks of (1,4)-linked β-Dmannuronic acid (M) and α-L-guluronic acid (G) residues. The blocks are composed of consecutive G-residues (GGGGGG), consecutive M-residues (MMMMMM), and alternating M and G-residues (GMGMGM) (Figure 19.1). The relative amount as well as length of each block depends on the origin of the alginate. The G-blocks of alginate can be intermolecularly cross-linked with divalent cations (e.g. Ca2+) through ionic interactions with the carboxylic groups in the alginate, inducing gelation. The composition, sequences, and molecular weight generally determine the physical properties of alginate (George and Abraham, 2006), as well as the phenotype of cells encapsulated in alginate gels. For example, the proliferation of rat bone marrow cells encapsulated into alginate gels containing a high G-block content was dependent on the composition (Wang et al., 2003).

19.2.2 Biocompatibility The biocompatibility of alginate has been extensively investigated and tested in vitro as well as in vivo, and yet remains controversial. Alginate can induce NaOOC

NaOOC O

OH

O

HO

OH

OH

O

HO O

OH O COONa

O

O OH O

OH G

COONa G

M

M

19.1 Chemical structure of sodium alginate. Alginate is composed of 1,4-linked β-D-mannuronic acid (M) and α-L-guluronic acid (G) residues arranged in a block-wise manner.

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a foreign body reaction and fibrosis, likely due to the toxicity of alginate (Cole et al., 1992; De Vos et al., 1996), but other studies have found little or no immunoresponse around alginate implants (Zimmermann et al., 1992). Alginates purified by multi-step chemical extraction procedures did not induce a significant foreign body reaction when implanted into animals (Klöck et al., 1994). Commercially available, purified alginates did not induce a foreign body reaction at least three weeks after implantation into animals (Mumper et al., 1994). The immunogenic response at the injection and implantation sites was attributed to toxic impurities in the alginate.

19.2.3 Mechanical properties Controlling the mechanical properties of hydrogels is critical in tissue engineering, as hydrogels need to maintain sufficient physical integrity until they are replaced by newly formed tissues (Wendt et al., 2005). The mechanical properties of gels can be controlled by physical and/or chemical methods. The physical properties of gels can generally be improved by increasing the molecular weight of a polymer, but alginate solution with a high molecular weight becomes increasingly viscous. Greater viscosity is often undesirable in processing (LeRoux et al., 1999), and cells mixed with the alginate solution risk damage from the high shear forces generated during mixing and injection (Kong et al., 2003). Manipulation of the molecular weight and its distribution, specifically formulated from a combination of high and low molecular weight alginates, can overcome this problem. Adjusting the molecular weight distribution with properly tailored alginate molecules can independently control the microstructures of the pre-gel solution and post-gel material. When this approach was applied, the elastic modulus of the gels increased significantly, while the viscosity of the pre-gel solution was only minimally raised (Kong et al., 2002). The compression modulus of alginate hydrogels ranges from less than 1 kPa to greater than 1000 kPa, while the shear modulus has values in the range of 0.02 – 40 kPa. In addition, these properties are highly dependent on the alginate composition, cross-linker type, and gelling and storage environments (Drury et al., 2004). The gelation temperature influences the mechanical properties of gels because it alters gelation time. At lower temperatures, the activity of ionic cross-linkers (e.g. Ca2+) is reduced and cross-linking becomes slower. The resulting cross-linked network structure has greater order, resulting in enhanced mechanical properties (Augst et al., 2006). The composition of the aqueous phase of an alginate solution also significantly influences the mechanical properties of the resultant gels. For example, phosphate ions in an alginate solution temporarily bind to calcium ions that cause slower gelation of the alginate solution. Alginate can be covalently cross-linked with various cross-linking molecules using

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carbodiimide chemistry, and the cross-linking density significantly influences the mechanical properties of the gels (Eiselt et al., 1999) (Figure 19.2). Interestingly, cells can provide additional mechanical integrity to alginate gels through interactions between cells and cell adhesion ligands coupled to alginate chains (Drury et al., 2005).

19.2.4 Degradation Alginate is inherently non-degradable in physiological conditions, but ionically cross-linked alginate gels can be dissolved via a process involving a loss of divalent ions into the surrounding media rather than actual degradation. In addition, the molecular weights of many commercially available alginates are typically above the renal clearance threshold of the kidney (Al-Shamkhani and Duncan, 1995). An attractive approach to controlling the degradation of alginate gels includes the use of oxidized guluronate blocks that are isolated from alginate. In brief, commercially available alginate is hydrolyzed under acidic conditions to break down glycosidic linkages between mannuronate and guluronate residues, after which polyguluronate (PG) can be isolated at pH 2.85 (Mw = 6000 Da). PG is then oxidized with sodium periodate to prepare poly(aldehyde guluronate) (PAG), which can be subsequently crosslinked with adipic acid dihydrazide (AAD) to form gels (Figure 19.3). The

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19.2 Changes in elastic moduli of alginate hydrogels covalently cross-linked with poly(ethylene glycol). Values represent mean (n = 5) and standard deviation. Reproduced with permission from Macromolecules 1999, 32, 5561–5566. Copyright 1999 American Chemical Society.

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19.3 Chemical structure of poly(aldehyde guluronate) hydrogel covalently cross-linked with adipic acid dihydrazide. Reproduced with permission from Macromolecules 2000, 33, 97–101. Copyright 2000 American Chemical Society.

resultant hydrazone bond between PAG and AAD is degradable, and the PAG gels are degradable in aqueous media by a bulk erosion process. The higher the AAD concentration used to form a gel, the slower the degradation rate. The degradation rate and mechanical properties are typically coupled, but gels with a high content of dangling single-end AAD molecules showed retarded degradation behavior, irrespective of the low cross-linking density (Lee et al., 2000a). Slightly oxidized alginate can be degradable in aqueous media, and has demonstrated a potential for delivery of chondrocytes in the engineering of cartilage (Bouhadir et al., 2001). The periodate oxidation cleaves the carboncarbon bond of the cis-diol group in the uronate residue and alters the chair conformation to an open-chain adduct, which significantly influences scission of the alginate backbone. A significant reduction in the molecular weight of the oxidized alginate was observed during incubation, and it was dependent on the pH and temperature of the medium. The oxidation of sodium alginate did not significantly interfere with the formation of ionic junctions with divalent cations, and gels formed in the presence of calcium ions. Degradable alginate gels containing chondrocytes were subcutaneously injected into mice, and the resultant construct after seven weeks of implantation was rigid and had a white opalescence, consistent with the appearance of native cartilage.

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Methods of gelling

19.3.1 Ionic cross-linking The gelation of alginate can be achieved under relatively mild conditions with an ionic cross-linking agent such as divalent cations (e.g., Ca2+). The cross-linking of alginate is mainly achieved by exchange of sodium ions from the guluronic residues with divalent cations, resulting in the formation of an egg-box structure (Figure 19.4). The divalent cations bind to the α-Lguluronate blocks in a highly cooperative manner (Smidsrod and SkjakBræek, 1990), and each alginate chain forms junctions with other chains to generate cross-linked network structures (Dupuy et al., 1994). Although calcium chloride (CaCl2) is one of the most frequently used ionic crosslinking molecules to prepare alginate gels, it typically forms gels very quickly. As a result, calcium sulfate (CaSO4) and calcium carbonate (CaCO3) are often used to fabricate alginate gels. For example, an alginate solution can be mixed with CaCO3, which is not soluble in water at neutral pH. Glucono-δ-lactone is then added to the alginate/CaCO3 solution in order to dissociate Ca2+ from the CaCO3 by lowering the pH. The released Ca2+ subsequently initiates the slow gelation of the alginate solution (Crow and Nelson, 2006). Although ionic cross-linking is a direct and simple method for alginate gel formation, the mechanical properties and stability of the resultant gels are very limited. Many approaches have been reported to fabricate stronger and more stable alginate gels cross-linked with divalent cations. Gelation time is an important factor in controlling gel uniformity and strength. Slower gelation produces more uniform structures and greater mechanical integrity (Kuo and Ma, 2001). The properties of ionically cross-linked alginate gels vary significantly depending on the chemical structure of alginate. The degradation behavior of ionically cross-linked gels can be regulated by controlling the dissociation rate of the polymer chains. Alginate gels with high M-block content (MVM) show a faster decline in the elastic modulus than gels with high G-block content (MVG), indicating that MVM

Ca2+ ( )

19.4 Alginate hydrogels can be formed by ionic cross-linking with calcium ions (egg-box model).

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gels undergo a more rapid loss of cross-links than MVG gels. The fast loss of cross-links within gels that possess geometrically mismatched cross-linked junctions significantly reduces the mechanical rigidity (Kong et al., 2004).

19.3.2 Covalent cross-linking Ionic cross-linking of alginate with calcium ions may cause high cytotoxicity and large foreign body reactions due to the excess amount of calcium ions in the gel. The calcium ions released from the gel through the exchange of sodium ions in the blood and wound exudates promote a hemostatic process, while the gel serves as a matrix for aggregation of platelets and erythrocytes (Suzuki et al., 1998). Covalent cross-linking has been widely investigated in efforts to improve the physical properties of gels for many applications, including tissue engineering. Covalent cross-linking creates gels with improved mechanical properties, compared with those of ionically cross-linked ones. However, many covalent cross-linking reagents are toxic, and the unreacted cross-linkers need to be thoroughly removed from gels. Covalent cross-linking of alginate with poly(ethylene glycol)-diamines of various molecular weights was investigated in order to prepare gels with a wide range of mechanical properties. The elastic modulus increased gradually with an increase in cross-linking density or weight fraction of poly(ethylene glycol) (PEG) in the gel (Eiselt et al., 1999). The mechanical properties and swelling of alginate hydrogels can be tightly regulated by using different kinds of cross-linking molecules and by controlling the cross-linking densities. The chemistry of the cross-linking molecules significantly influences hydrogel swelling. The introduction of hydrophilic cross-linking molecules as a second macromolecule (e.g. PEG) compensates for the loss of hydrophilic character of the hydrogel during the cross-linking reaction between hydrophilic groups in the alginate backbone (Lee et al., 2000b). Porous alginate gels were prepared by covalent cross-linking with ethylenediamine and used for tissue engineering applications. Cytotoxicity tests demonstrated that these gels were more biocompatible than alginate gels cross-linked with calcium ions. There was also less foreign body reaction caused by the covalently cross-linked gels in pigs, consistent with the results of in vitro cytotoxicity tests (Suzuki et al., 1998).

19.3.3 Photo cross-linking Photo cross-linking is an exciting approach to in situ gelation that can provides improved mechanical strength compared to ionic cross-linking. Alginate modified with methacrylate and cross-linked with laser exposure (argon ion laser, 514 nm) for 30 s in the presence of eosin and triethanolamine forms a

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clear, soft, and flexible hydrogel. The gel was useful for sealing corneal perforation in vivo, indicating a potential clinical use for sutureless surgery (Smeds and Grinstaff, 2001). Photo cross-linking reactions involve the use of a light sensitizer or the release of acid, which may be harmful to the body. The alternative photo cross-linking approach uses poly(allylamine) partially modified with α-phenoxycinnamyldiene acetylchloride, which dimerizes upon light exposure at about 330 nm and releases no toxic byproducts during the cross-linking reaction (Tanaka and Sato, 1972). The mechanical properties of the gels formed from this photosensitive poly(allylamine) and alginate were significantly improved by light irradiation, and the gels were freely permeable to cytochrome c and myoglobin, but less permeable to serum albumin (Lu et al., 2000).

19.3.4 Cell cross-linking A number of approaches have been reported to promote gel formation through ionic, covalent, and photo cross-linking, and temperature-dependent phase transition. However, the ability of cells to contribute to gel formation has been largely ignored. Alginate can be cross-linked with cells via specific receptor-ligand interactions without using any additional cross-linking molecules. Cell-interactive alginate was synthesized by chemically coupling cellular adhesion ligands (e.g. a peptide with the sequences of arginineglycine-aspartic acid, RGD) using water-soluble carbodiimide chemistry. The existence of the ligand dramatically influenced the distribution of cells in the alginate solution, and subsequently generated the cross-linked network structure to form gels (Lee et al., 2003). Cells added to the RGD-modified alginate solution form a uniform dispersion within the solution, due to the ability of integrin receptors on the cell surface to bind and cross-link multiple polymer chains of the RGD-modified alginate (Figure 19.5). In contrast, cells added to unmodified alginate solutions aggregate to form a non-uniform structure, owing to negligible cell-polymer interactions and dominant cellcell interactions in the system. The RGD-modified alginate solution formed gels in the presence of cells within 600 s, and this gelation behavior was repeated following multiple shear-induced breakdowns of the gels, indicating reversible gelation behavior of the system. It was considered ideal as an injectable cell delivery system because a gel can flow as a liquid when injected but re-gel once placed in body. In addition, adhesive interactions between cells and ligands coupled to the gel-forming materials enhance the overall mechanical properties of the gels. Above the critical ligand and cell density, C2C12 myoblasts encapsulated in RGD-modified alginate gels provided additional mechanical integrity to the gels by adding cell-ligand cross-linking to the usual ionic cross-linking (Drury et al., 2005).

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19.5 Photomicrographs of (a) RGD-modified alginate/cell mixture and (b) unmodified alginate/cell mixture. Enlarged schematics demonstrate the concept of network formation by specific interactions between cells and polymer chains. Reproduced with permission from Advanced Materials 2003, 15, 1828–1832. Copyright 2003 Wiley-VCH.

19.4

Applications of alginate hydrogels in tissue engineering

19.4.1 Cartilage Tissue engineering approaches using three-dimensional scaffolds, including hydrogels, have proved useful for transferring autologous chondrogenic cells and restoring damaged cartilage. Alginate gels have been frequently used to engineer cartilage. Suspensions of chondrocytes in alginate solution were mixed with calcium sulfate and injected into molds of facial implants in order to engineer pre-shaped cartilage. These constructs were subcutaneously implanted into mice and sheep. After 30 weeks of implantation, the constructs formed cartilage with three-dimensional shape retention, and the proteoglycan and collagen contents of the constructs reached approximately 80% of the values for native cartilage. The equilibrium modulus and the hydraulic permeability were 74 and 105% of those of native auricular cartilage, respectively (Chang et al., 2001; Chang et al., 2003). Macroporous alginate gels with predefined geometries were prepared, compressed into significantly smaller, temporarily dried forms, and introduced into mice through a small catheter. The gels were then rehydrated in situ with a suspension of primary bovine articular chondrocytes and recovered their original shape and size within 1 h, which allowed cartilage formation in the animals with the desired geometry. This approach to engineering tissues with shape-memorizing scaffolds may enable the minimally invasive delivery of cells and the

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formation of new tissues with desired shapes and sizes in vivo (Thornton et al., 2004). Human mesenchymal stem cells (MSCs) were encapsulated in alginate gels and cultured in serum-free medium with the addition of transforming growth factor (TGF)-β1, dexamethasone, and ascorbate 2-phosphate. MSCs in alginate gels formed cartilage, and this system represents a relevant model for the study of the molecular mechanisms involved in the chondrogenesis and endochondral ossification (Ma et al., 2003). The chondrogenic differentiation of human adipose-derived adult stem cells seeded in alginate gels, as well as the functional properties of the engineered cartilage, were significantly enhanced compared to control conditions (Awad et al., 2004).

19.4.2 Bone Modification of alginate with an RGD-containing peptide promoted osteoblast adhesion and spreading, which were dependent on the adhesion ligand density. MC3T3-E1 cells demonstrated increased osteoblast differentiation with the peptide-modified hydrogels, which was confirmed by the up-regulation of bone-specific differentiation markers. Further, transplantation of primary rat calvarial osteoblasts into mice using RGD-modified alginate gels enhanced in vivo bone formation (Alsberg et al., 2001). Chondrocytes and osteoblasts were co-transplanted into mice using the RGD-modified alginate hydrogels, and the transplanted cells were organized into structures that morphologically and functionally resembled growth plates. New bone tissue was formed that grew in mass and cellularity by endochondral ossification in a manner similar to normal long-bone growth (Alsberg et al., 2002). Degradable and injectable alginate-derived hydrogels, composed of PAG and AAD, were prepared, mixed with rat primary calvarial osteoblasts, and subcutaneously injected into the backs of mice. Mineralized bone tissues were observed after 9 weeks of implantation (Figure 19.6) (Lee et al., 2001). Various alginate/hydroxyapatite (HAP) composite scaffolds were prepared by a phase separation method, containing a well-interconnected porous structure with an average pore size of 150 µm and over 80% porosity. Rat osteosarcoma UMR106 cells, seeded into the scaffold, showed better cell adhesion to alginate/HAP composite scaffolds than to pure alginate scaffolds (Lin and Yeh, 2004). Cell-encapsulating alginate gel beads were introduced into calcium phosphate cement, where the alginate beads adequately protected the cells and resulted in favorable cell viability, indicating a potential for bone tissue engineering in moderate stress-bearing applications (Weir et al., 2006). In contrast to rat cells, human cells did not readily attach or proliferate on alginate gels. However, gels containing both collagen type I and β-tricalcium phosphate enhanced human cell adhesion and proliferation, indicating a potential use of alginate gels for human tissue engineering (Lawson et al., 2004).

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19.6 Photomicrograph of tissue section where osteoblasts were transplanted using poly (aldehyde guluronate) gels cross-linked with adipic acid dihydrazide. Tissue section was taken after nine weeks of implantation and stained with the von Kossa method to highlight mineralized tissue. Labels indicate the remaining gel (H) and newly formed bone tissue (B). Reproduced with permission from Journal of Biomedical Materials Research 2001, 56, 228–233. Copyright 2001 Wiley-VCH.

Three-dimensional cell culture systems were prepared from calcium-crosslinked alginate gels and used to investigate proliferation and differentiation of rat bone marrow cells. Gel thickness was critical in determining cell behavior, and different geometries did not influence cell differentiation (Barralet et al., 2005). Bone marrow stromal cells were isolated, expanded, and induced into osteogenic cells in a defined medium in vitro, and then mixed with calcium cross-linked alginate gels in order to repair horizontal alveolar bone defects in dogs. Bone nodule structure was observed at four weeks postsurgery and the engineered bone became more mature after 12 weeks, which was similar to normal bone (Weng et al., 2006). Alginate/chitosan gels containing mesenchymal stem cells and bone morphogenetic protein-2 also show a potential for new bone formation in animals, especially trabecular bone formation in mice (Park et al., 2005).

19.4.3 Nerve Alginate gel can be used as glue for artificial nerve guides for peripheral nerves and for repair of disrupted pathways in central nervous tissues that cannot be sutured (Suzuki et al., 2000). Alginate-based highly anisotropic

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capillary hydrogels (ACH) were introduced into acute cervical spinal cord lesions in adult rats and integrated into the spinal cord parenchyma without major inflammatory responses. They also maintained their anisotropic structure, indicating that alginate-based gels can provide a promising strategy to induce directed nerve regrowth following spinal cord injury (Prang et al., 2006). Alginate gels have also been used for peripheral nerve regeneration. A viscous injectable alginate solution exhibited excellent biocompatibility and strong inhibition of perineurial granulation in a rat neurolysis model. The alginate solution was absorbed completely within six weeks without inducing inflammation, and enhanced repair of the perineurium through regeneration of epithelial-like cell layers (Ohsumi et al., 2005). Covalently cross-linked alginate scaffolds enhanced nerve regeneration in peripheral nerves and spinal cords. Alginate gels, covalently cross-linked with ethylenediamine, were implanted into cat sciatic nerves. Remarkable restoration of a 50-mm gap in cat sciatic nerves was obtained using alginate gels (Hashimoto et al., 2005). Covalently cross-linked alginate scaffolds promoted the outgrowth of regenerating axons and astrocyte reactions at the stump of transected spinal cords in young rats (Kataoka et al., 2004). Nerve conduits (NCs) made of a hydrogel consisting of alginate and chitosan have also been developed as a promising alternative to conventional treatments for peripheral nerve repair. The NCs possess sufficient mechanical strength and remarkable tear resistance (Pfister et al., 2007). Hydrogels with tubular microstructures were prepared from self-assembled copper-capillary alginate gels and treated with oligochitosan to create stable scaffolds through polyelectrolyte complex formation. Mouse embryonic stem cells were guided to form cylindrical structures within scaffold capillaries, suggesting a unique and potential platform of alginate gels for stem cell-based tissue engineering (Willenberg et al., 2006). Mouse-derived neural stem cells (NSCs) with the capacity of extensive self-renewal and multilineage differentiation were cultured in three-dimensional calcium alginate beads, and expanded within the gel (Li et al., 2006). The alginate gel surface was modified with a peptide with the sequence of YIGSR, and proved useful to promote the adhesion of NB2a neuroblastoma cells and neurite outgrowth from these cells, depending on the peptide density on the gel surface (Dhoot et al., 2004).

19.4.4 Liver Insufficient donor organs for liver transplantation have increased the requirement for new therapies for liver disease, and tissue engineering is a promising approach to developing hepatic tissues suitable for the functional replacement of a failing liver. A potential approach to enhance the performance of implanted hepatocytes includes their aggregation and re-expression of their differentiated function prior to implantation using alginate scaffolds

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(Wang et al., 2003; Selden and Hodgson, 2004). The hydrophilic nature of the alginate scaffold as well as its porous structure and interconnectivity enable the efficient seeding of hepatocytes into the scaffolds. In high density cellular constructs, hepatocellular functions remained high (Dvir-Ginzberg et al., 2003). Freshly isolated rat adult hepatocytes seeded in alginate scaffold did not proliferate over two weeks, but nearly all the seeded cells maintained viability, and their aggregation was enhanced. The cells appeared to synthesize fibronectin, which was deposited on the spheroids and promoted their functional expression in the scaffold (Glicklis et al., 2000). Culture of C3A cells (a human hepatocyte cell line) within alginate scaffolds induced the formation of spheroids that displayed multilayer cell morphology and were characterized by a large number of tight junctions, polar cells, and bile canaliculi, making them similar to spheroids of primary hepatocytes (Elkayam et al., 2006). A highly porous scaffold composed of alginate and galactosylated chitosan was fabricated for developing bioartificial liver devices. Hepatocytes were aggregated to form multicellular spheroids in the scaffold, and intercellular molecules such as connexin32 and E-cadherin genes related to cell-cell interactions were expressed. In addition, co-culture of hepatocytes with NIH 3T3 resulted in enhanced liver-specific functions, such as albumin secretion rates (Seo et al., 2006). The porous structure of alginate scaffolds was controlled by the freeze-dry method. Rat hepatocytes seeded within the scaffold were distributed according to the pore shape. For example, cells formed lines and spheroid-like aggregates in the pores with elongated shape and isotropic spherical shape, respectively, indicating that pore shape can modulate hepatocyte morphogenesis (Zmora et al., 2002).

19.4.5 Ovarian follicle In vitro systems for follicle culture have been developed to preserve fertility for women, as oocytes grown in vitro are of low quality and give few live births. Immature mouse follicles were cultured in vitro using an alginate hydrogel to maintain the three-dimensional architecture of oocytes and cellcell interactions. Follicles encapsulated into the gel developed mature oocytes with the capability of fertilization, similar to that of oocytes matured in vivo (Xu et al., 2006). Alginate hydrogels modified with an ECM component (e.g. laminin) or RGD peptide promoted follicle maturation and produced meiotically competent oocytes, indicating that the ECM is a dynamic regulator of follicle development (Kreeger et al., 2006). Ovarian follicles were also encapsulated into calcium-cross-linked alginate gels and cultured in vitro, in order to support three-dimensional follicle architecture while maintaining responsiveness to follicle stimulating hormone (FSH) stimulation. Threedimensional follicle growth was achieved, and the inclusion of FSH in the alginate gels restored follicle growth in response to FSH (Heise et al., 2005).

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This approach may be useful to human egg banks for fertility preservation in women and endangered species.

19.5

Summary and future trends

Alginate has demonstrated potential as a biomaterial for many biomedical applications, particularly tissue engineering. The most attractive features of alginate in tissue engineering include low toxicity and ease of gelation. A number of methods to control gelling behavior of alginate solutions and the physical and chemical properties of the resultant gels have been reported to date. The introduction of appropriate cell adhesion ligands to alginate gels and their spatial organization in the gels will be a challenge in tissue engineering, as this strategy can regulate specific interactions between cells and alginate gels three-dimensionally, making it more biologically relevant to living organisms. In addition, delivery of soluble factors such as growth factors and DNA using alginate gels is critical to creating functional and clinically successful tissues. Localized delivery of soluble factors (single or multiple) using alginate gels can protect the factors from proteolytic degradation in the body, and release of the factors from the gels in a sustained manner can help promote tissue regeneration.

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References

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20 Fibrin matrices in tissue engineering B. T A W I L, H. D U O N G and B. W U, University of California Los Angeles, USA

20.1

Introduction

Tissue engineering scaffolds present structural and biological cues to guide cell attachment, differentiation and function. The microarchitecture of an ideal scaffold should facilitate transport of gases, nutrients, peptides and macromolecules needed for cellular activities and tissue development. Scaffold materials must be non-toxic and non-immunogenic; be mechanically compatible with the biological environment; and undergo biodegradation at the appropriate time. From a practical perspective, the materials should be readily available, easy to produce reliably and reproducibly, and well recognized by regulatory agencies due to extensive documentation in human use. Among the multitude of substances that are currently being used for scaffolds, fibrin appears to be a suitable scaffold material that satisfies many of the above criteria in a wide range of biological indications. Fibrin is nature’s scaffold following tissue injury to initiate hemostasis, and provides the initial matrix into which cells invade to rebuild the damaged tissue. Fibrin facilitates cell adherence, migration and biochemical interactions. As cells invade and proliferate inside the fibrin scaffold, they secrete proteases that break down fibrin. These cells also secrete specific extracellular matrix molecules, such as collagen, to remodel the damaged tissues. The availability and relative ease of isolating and purifying fibrinogen and thrombin, the key coagulation factors that form the fibrin polyfibrils, make fibrin an attractive material for use in tissue engineering. The fibrin dimeric molecule has been well elucidated (Boyles et al., 1950; Horowitz et al., 1957; Hall and Slayter, 1958; Budzynski and Olexa, 1981), and there has been increased interest in its wide-ranging applications (Meana et al., 1998; Mah-Becherel et al., 2002; Ho et al., 2006). Fibrin has been studied in the treatment of chronic wounds, bone defects, Parkinson and Alzheimer patients and in Muscular Dystrophy, etc. by delivering growth factors and cells. In this chapter, we will discuss how fibrin is used as a scaffold in generating tissue-engineering products. 533 © 2008, Woodhead Publishing Limited

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20.2

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Fibrin formation

Following injury, fibrin clot is formed to prevent further blood loss. Clot formation is achieved by activating platelets and a complicated cascade of coagulation factors that result in the activation of thrombin, which proteolytically cleaves fibrinogen dimers to form fibrin monomers, (α, β and γ chains). The activated fibrin monomers spontaneously polymerize into loosely bound fibrin fibrils. Subsequently, Factor XIIIa further stabilizes the fibrin fibrils to form an insoluble fibrin clot (Amrani et al., 2001; Ariens et al., 2002). Calcium and vitamin K are essential cofactors in activating the various modulators. During fibrin clot formation, platelets, growth factors such as FGF and VEGF, and extracellular matrix (ECM) proteins such as fibronectin are trapped within the fibrin clot (Diegelmann and Evans, 2004; Werner and Grose, 2003), thereby, creating a rich environment for cell proliferation and differentiation. While the initial function of the fibrin clot is to cease bleeding, its longterm function is to act as a scaffold for the various invading cells that clean and rebuild the wounded area (Clark, 1998; Tawil et al., 2005). First, monocytes invade the fibrin and get activated to macrophages which cleanse the wound site of any cell debris or bacterial contamination. A few days later, fibroblasts, keratinocytes and endothelial cells migrate into the fibrin scaffold, proliferate and secrete growth factors, proteases, cytokines, and ECM proteins such as collagen to rebuild the dermis, epidermis and revascularize the wounded area respectively (Werner and Grose, 2003; Diegelmann and Evans, 2004). Fibrin is also formed during injury to other types of tissues involving different types of cells invading the fibrin scaffold. In bone injury, for example, osteoblasts invade the fibrin to rebuild the bone. The important role of fibrin as a scaffold for invading cells and its inflence on wound healing is clearly shown in a fibrinogen deficient mice (Drew et al., 2001); while the time needed to heal was similar in both the fibrinogen knockout and control mice, there was an altered cell migration in the knockout mice that afffected the wound breaking strength (Drew et al., 2001). Structural features such as fibrin porosity, number of branch points, and fibril thickness can reportedly affect wound healing (deMaat et al., 2006). Furthermore, the fibrinogen molecular weight may influence angiogenesis (deMatt et al., 2006). Thus, fibrin formation at the onset of an injury is essential for a timely and complete healing of an injured tissue. The role of fibrin during the tissue remodeling process is biphasic: first by ceasing the bleeding initially and second, by functioning as a structural support for cell growth, invasion, attachment and biochemical activity.

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Fibrin use in surgery

The major commercial biomedical application of fibrin is its use as haemostatic agent during surgery (Pfluger et al., 1986; Salvino et al., 1993; Radosevich et al., 1997) by formulating hyper-physiological concentrations of fibrinogen and thrombin to the point where the hemostasis occurs instantaneously (Kjaergard et al., 1994). Fibrinogen concentration in the body is normally 2– 5 mg/ml; however, in the commercially available products the concentration could reach 100 mg/ml. The available products differ in fibrinogen content as well as the amount of fibrinolysis inhibitor such as Aprotinin, which increases the half-life of the fibrin clot especially in a protease-rich environment (Despotis et al., 1995). Derived from virus-inactivated human plasma, fibrin is an important biomaterial in surgery (Tawil, 2005). Fibrin is approved for various surgical procedures including cardiopulmonary bypass surgery, trauma to the spleen and sealing of colostomy closure (Codispoti and Mankad, 2002; Busuttil, 2003). Fibrin is also used in a laparoscopic gastric bypass in the prevention of anastomotic leak and internal hernia (Silecchia G, et al., 2006), in the control of aerostasis with a high risk of air-leakage (Massone et al., 2003), in the control of the intraoperative cerebrospinal fluid-leakage during trans-nasal/trans-sphenoidal pituitary microsurgery (Sade et al., 2006) and to improve healing in the oral cavity (Yucel et al., 2002) and to reduce the incidence of lymphatic complications during the closure of inguinal wounds (Giovannacci et al., 2003). Other applications for fibrin include: usage in cosmetic surgeries, i.e. facial rejuvenation procedures (Matarasso et al., 2005; Marchac and Greensmith, 2005), and as a substitute for sutures during conjunctival autograft surgery for primary pterygium (Marticorena et al., 2006). Furthermore, a study showed that injecting 1–2 ml of fibrin in the seam of a meniscus repair enhanced healing of the tissue (Henning et al., 1990).

20.4

Fibrin matrices to deliver bioactive molecules

During hemostasis, the fibrin clot stabilizes the platelet plug and mediates the release of numerous growth factors by the platelets, such as PDGF, FGF, VEGF, TGF-β, etc. Not surprisingly, fibrin as a scaffold has been used to deliver growth factors and active peptides. Some growth factors bind either weakly (MacPhee et al., 1996; Wong et al., 2003) or strongly (Sahni et al., 2000; Hasimoto et al., 1992; Albes et al., 1994; Wong et al., 2003) to fibrin scaffold. A study that examined the release kinetics of basic Fibroblast Growth Factor (bFGF) and two forms of Vascular Endothelial Growth Factor (VEGF165 and VEGF121) added to fibrin clots (Wong et al., 2003), found that bFGF was more slowly released than VEGF165 and VEGF121. Furthermore, using chorioallantoic membrane (CAM) model of neovascularization, these growth

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factors demonstrated angiogenic activity. Other studies have shown that some growth factors could be cross-linked to fibrin and released slowly as fibrin degrades over a period of time (Andree et al., 2001; Schense et al., 2000). Fibrin is also used to deliver active peptides such as Macrophage Activator Lipoprotein (MALP-2), which is normally secreted by mycoplasma and which plays an important role in wound healing (Cole et al., 2007). In Cole et al. (2007), it was shown that monocytes seeded on fibrin containing MALP-2 secreted more cytokines such as IL-6, TNF-alpha, and chemoattractants such as MIP -1 alpha and MCP-1 when compared to monocytes seeded on a 3D fibrin clot in the absence of MALP-2. These studies clearly show the potential of using fibrin to deliver growth factors, active peptides, painkillers, and antibiotics for the treatment of various diseases. Companies such as Kuros are already using fibrin to deliver bioactive substances such as parathyroid hormone peptide (PTH(1-34)) for the treatment of subchondral cystic lesion (Fuerst et al., 2007) and in the treatment of venous ulcers (Dragieva et al., 2004).

20.5

Fibrin – cell constructs

As indicated above, fibrin is used in surgery to stop bleeding; however, it is also a 3D scaffold that could be used to deliver cells to treat various diseases. What makes fibrin an ideal cell delivery vehicle is the ability to modulate its mechanical and biochemical characteristics. One major approach to modulate the 3D fibrin scaffold is to change the initial fibrinogen and thrombin concentration of the final fibrin scaffold as shown in Ho et al. (2006). The study reported that the microstructure involving the meshwork of the fibrin fibril correlates to the starting fibrinogen and thrombin concentrations. A scaffold consisting of a lower fibrinogen and thrombin concentration formed more open, homogeneous microstructures; whereas, a scaffold consists of a higher fibrinogen and thrombin concentration formed dense and nonhomologous microstructures (Ho et al., 2006). Here, we present some data to show fibrin characteristics in association with delivering different types of cells.

20.5.1 Fibroblasts in 3D fibrin scaffold We have studied the behavior of human dermal fibroblasts (proliferation and migration) in different fibrin scaffolds as a consequence of changing the initial fibrinogen and thrombin concentrations in the final fibrin scaffold (Cox et al., 2004). Fibroblasts proliferated robustly in a 3D fibrin scaffold that consists of 5 to 17 mg/mL fibrinogen and 1 to 167 U/mL thrombin concentrations. Fibroblasts also started to migrate out of the fibrin scaffold

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after a few days again depending on the final fibrin composition, i.e. fibrinogen and thrombin concentration. The expression of cytokines and growth factors was also dependent on the fibrin composition. Similarly, gene expression of cells seeded on 3D was dependent on fibrin composition (Coles et al., 2001). The modulation of gene expression was also observed in fibroblasts incorporated in a collagen gel (Lambert et al., 2001). The effect of cell proliferation, and gene and protein expression by fibrin composition indicates that there is a biochemical interaction between the cells embedded in the fibrin and the fibrin itself. The fibrin–fibroblasts constructs were further tested in animal studies (Mogford et al., 2004). Fibrin– fibroblasts constructs in the rabbit ear cutaneous wound model showed an increased granulation tissue area within the wounds vs. fibrin alone. Adding PDGF-B to the fibroblast–fibrin constructs further increased the granulation tissue. Other studies have shown that fibrin down-regulates the expression of ECM molecules and growth factors such as TGFβ-1 in fibroblasts obtained from adhesion surgery sites, thereby reducing the postoperative adhesion sites (Saed et al., 2004). Preliminary data obtained in our lab showed that the mechanical characteristics such as firmness and porosity of a 3D fibrin scaffold correspond to the fibrinogen and thrombin concentrations (Tawil et al., Preliminary Data). Scaffold stiffness and strength increased with an increasing concentration of fibrinogen and thrombin. The addition of cells to the fibrin scaffold further influences its mechanical properties. Since fibroblasts proliferate inside the 3D fibrin scaffold and form a rich meshwork of cell extensions, they might be exerting intrinsic forces upon the scaffold, thereby increasing the stiffness of the supporting scaffold. Conversely, these proliferating cells can also affect the scaffold’s mechanical properties by secreting ECM proteases that would breakdown the supporting fibrin fibril structure of the scaffold (Tuan et al., 1996). However, the same cells can secrete and synthesize another matrix molecule, collagen, which can, in turn, replace the structural support lost due to fibrin degradation. (Miller et al., 2005). The dynamic steady state between degradation and reconstruction of the fibrin scaffold by cellular interactions remains to be elucidated further. For the treatment of chronic wounds, a company, Intercytex in the UK, is using fibrin to deliver allogenic fibroblasts for the treatment of chronic wounds. ICX-PRO® is in Phase III clinical trials. It is manufactured in one day and can be stored for 21 days in a simple refrigerator (Kemp, 2006). The future work is presently focusing on how cell types, cell morphology, and cell density can alter the mechanical properties of the fibrin scaffold. Further studies will focus on the modulation of mechanical characteristics of fibrin scaffolds due to degradation of fibrin, synthesis of new matrix materials, cellular morphological changes, and other biochemical responses characteristic to the cell types.

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20.5.2 Monocytes in 3D fibrin scaffold Monocytes play a major role in the wound healing process, and they are the first cells to invade the fibrin clot (Clark, 1998). They play a role in clearing out the wounded area of any bacterial contamination or cell debris. Studies have shown that monocytes adhere and proliferate differently on fibrin depending on its fibrinogen and thrombin composition (Mana et al., 2006). Monocytes also migrate through the 3D-fibrin clots in 1–2 days (Mana et al., 2006). Similarly to fibroblasts, the protein expression in monocytes seeded on 3D-fibrin scaffold varied depending on the fibrinogen and thrombin concentrations (Mana et al., 2006). Monocytes interact with fibrin through their integrins (Simon et al., 1993; Perez and Roman 1995). Clustering of monocytes was also dependent on fibrin composition of the 3D scaffold (Mana et al., 2006). Neutrophil behavior in a 3D-fibrin clot is also dependent on the fibrin composition (Hanson and Quinn, 2002). These data indicate that fibrin is essential for the wound healing process, and that adding a proper fomulation of exogenous fibrin could improve tissue repair. Currently, there are no available products on the market to facilitate the delivery of monocytes for the enhancement of wound repair; nevertheless, it is a plausible approach that needs to be further explored.

20.5.3 Keratinocytes in fibrin scaffold As with fibroblasts embedded in fibrin scaffolds, recent studies showed that keratinocyte proliferation depended on the fibrin composition (Sese et al., 2003). High fibrinogen and thrombin concentrations in the final fibrin scaffold present the optimal environment for keratinocyte proliferation (Sese et al., 2003). Interestingly, the fibrin composition that favors fibroblast proliferation is not the same as the composition that favors keratinocyte proliferation supporting the model that there is a cross-talk between cells and their environment, i.e. fibrin. This suggest that different cell types may favor different fibrin compositions. Fibrin scaffolds have been used to co-culture multiple cell types to study cell to cell interactions. Sese et al. (2004) showed that there is cross-talk between fibroblasts and keratinocytes when mixed in a fibrin scaffold. They reported that fibroblasts cultured in scaffolds of specific fibrin compositions that are not conducive to their proliferation, proliferate robustly in these same compositions when co-culturing with keratinocytes. These data suggest that cell to cell interactions could enhance cell proliferation in an environment that is not favorable for growth, i.e high fibrin composition scaffolds. Fibrin has been used as a cell delivery vehicle such as the delivery of autologous keratinocytes for the treatment of chronic wounds. BioTissue, a German company, market a product called BioSeedR that involves the isolation

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of autologous keratinocytes from a biopsy of a chronic wound patient, propagating the cells for a few days, adding to the thrombin, and delivering to the wound site with the addition of fibrinogen (Horch et al., 2001).

20.5.4 Neuronal stem cells in fibrin scaffold Neuronal Stem Cells (NSCs) were also shown to proliferate in 3D fibrin scaffold differentially depending on the fibrin composition (Ferguson et al., 2003). Furthermore, their morphology depended on the fibrin composition and for how long the cells were embedded in the fibrin, i.e. they extended axons in some formulations (Ferguson et al., 2003). The potential of using fibrin to deliver neuronal stem cells for the treatment of neuronal diseases such as Alzheimer or Parkinson diseases should be further explored.

20.5.5 Mesenchymal stem cells in fibrin scaffold Mesenchymal stem cells (MSCs) are pluripotent progenitor cells; they can be stimulated in vitro to undergo differentiation into various other cell types such as adipose, chondrocyte and osteocyte. Recent improvement in the techniques for isolation and induction of differentiation of MSCs have allowed researchers to engineer tissues such as cartilages and bone. They have also begun to examine closely the mesenchymal stem cell–fibrin interaction and differentiation. For example, in Park et al. (2004), tissue-engineered cartilage was achieved using MSCs grown in fibrin/hyaluran composite. MSC’s morphology, proliferation, and osteogenic differentiation in 3D fibrin scaffold were also examined (Ho et al., 2006; Catelas et al., 2006). Fibrin scaffold containing a low concentration of fibrinogen presented an excellent environment for the cells to proliferate. However, osteogenic marker expression such as the enzyme alkaline phosphatase and nodules of mineralization, were stronger in fibrin compositions with a high fibrinogen concentration (Ho et al., 2006; Catelas et al., 2006). The use of fibrin to deliver mesenchymal stem cells for the treatment of bone will be most likely combined with the addition of synthetic polymers. While the synthetic polymers provide the load-bearing characteristic of such a construct, fibrin provides the optimal environment for mesenchymal stem cells to proliferate and differentiate. Fibrin (PEGylated) – mesenchymal stem cell construct could also be used to treat infracted heat (Zhang et al., 2006).

20.5.6 Endothelial cells in fibrin scaffold Tissue remodeling requires not only participation of fibroblasts but also other cell types. In response to the tissue damage, endothelial cells are stimulated

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and activated to migrate into the wound matrix where they undergo angiogenesis to provide vascular growth and remodeling. Proliferation, migration and invasion of surrounding tissues are essential characteristics of endothelial cells for the regeneration of new blood vessels. Numerous studies have shown that collagen matrices provide a good substrate for endothelial cell proliferation and migration (Scherberich et al., 2000; Sweeny et al., 1998; Sweeny et al., 2003). When embedded in fibrin scaffold, endothelial cells have been shown to produce vessel segments in the form of sproutings and tubule formations (Vailhe et al., 1998; Lafleur et al., 2002). Others have utilized fibrin scaffolds and endothelial cells to generate cardiovascular structures like vessels and valve conduits (Jockenhoevel et al., 2001). Clearly, the study of endothelial cell interactions within fibrin or other scaffold materials may give rise to invaluable findings associated with vascular biology. As for other cell types, could fibrin alone or fibrin with endothelial cells improve on the resolution of chronic wounds? Could the combination of cells, including fibroblasts, keratinocytes and endothelial cells, be the best approach? What is the cost associated with isolating and growing endothelial cells? While the use of fibroblasts and keratinocytes in the treatment of chronic wounds is a common strategy by various companies, the use of endothelial cells is under-examined and under-explored.

20.5.7 Fibrin and other cell types Other studies examined the use of fibrin to deliver urothelial cells onto a prefabricated pouch (Wechselberger et al., 1998). Fibrin was also shown to influence the proliferation and migration of human chondrocytes (Kirilak et al., 2006) as well as to contribute to the formation of a healthy neocartilage from chondrocytes obtained from older patients (Mesa et al., 2006). Fibrin, used to deliver skeletal myoblasts to treat the infarcted area of the left ventricle of a rat, showed a better outcome (in five weeks) than injecting the myoblasts alone (Christman et al., 2004). Furthermore, myofibroblasts, retrieved from human aortic tissue, proliferated in fibrin scaffold and produced collagen with no toxic degradation or inflammatory reactions (Ye et al., 2000).

20.6

Mechanical characteristics of fibrin scaffold

In order to use fibrin properly as a scaffold in various applications, one should appreciate the strength and limitations of such scaffolds. Many factors affect the fibrin scaffold’s mechanical characteristics such as fibril thickness, homogeneity, and porosity of the scaffold. Studies have shown that salt concentration influences the fibrin porosity (Yang et al., 2007). The elasticity of the fibrin scaffold is influenced by Factor XIII (Roberts et al., 1973;

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Mosesson et al., 2001). The fibrinogen concentration influences the breaking strength and adhesiveness of the fibrin scaffold (Sierra, 1993; Khare et al., 1998). Fibrin mechanical characteristics could further be manipulated by mixing it with synthetic materials. For example, fibrin mixed with Macroporous Biphasic Calcium Phosphate (MBCG) affected the overall fibrin structure (Daculsi et al., 2004). Fibrin could also be mixed with natural materials like the ECM glycoproteins: collagen, fibronectin, laminin, elastin and glycosaminoglycans. A study showed that fibrin mixed with collagen-derived sponges demonstrated an improved tissue-engineering scaffold (Yang et al., 2007). Presently, we are examining the mechanical characteristics of fibrin scaffolds formed with different fibrinogen and thrombin concentrations, fibrin-collagen scaffolds, and fibrin scaffolds containing different types of cells including either fibroblasts, keratinocytes or mesenchymal stem cells (Young et al., 2007). Our preliminary data show that the constructs’ stiffness corresponded to the relative concentration of fibrinogen to thrombin and that cells also have an influence on the overall stiffness of the fibrin scaffold.

20.7

Future trends

Scientific findings over the last 30 years has clearly supported the use of fibrin as a delivery vehicle for bioactive substances and various cell types. The potential for fibrin in an engineering technology that could be used alone or along with other technologies to treat various diseases is tremendous. Here we list the potential use of fibrin in future tissue-engineering products.

20.7.1 In cardiovascular Fibrin could be used alone or in association with cells to treat various cardiovascular diseases. For example, Randy Lee (Christman et al., 2004a,b) have shown that fibrin alone or fibrin with skeletal myoblasts increases the wall thickness, reduces infarct expansion, and induces neovasculature formation in a rat Myocardial Infarction model (Christman et al., 2004a,b). Fibrin could also be used to create pulsatile cardiac sheets that functionally integrate with the host, restoring its function without creating a serious arrhythmia (Furuta et al., 2006). Future use of fibrin could include its use to enhance angiogenesis in patients with peripheral arterial disease (PAD) which leads to a high mortality (Eberhardt and Coffman, 2004). PAD is caused by atherosclerosis of the lower extremities that leads to ischemia and consequently affects the healing process in that area. In addition to its use as a cell-carrier to deliver fibroblasts or keratinocytes to chronic wound, fibrin could deliver endothelial cells in the same patient to improve the blood circulation by rebuilding the blood vessels.

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20.7.2 In cartilage Surgeons used fibrin to secure the periosteal flap under which they transplanted autologous chondrocytes in the treatment of focal cartilage defects (Sohn et al., 2002). However, fibrin is also used to deliver chondrocytes to produce a healthy cartilage (Silverman, et al., 1999). This delivery could be done by entrapment of chondrocytes in fibrin–alginate beads creating a stable transplant (Perka et al., 2001). Fibrin could also serve as a carrier for growth factors such as TGFβ1, attracting mesenchymal stem cells and leading to a chondrogenesis (Huang et al., 2002).

20.7.3 In bone The use of fibrin in bone tissue-engineering is widespread due to the fact that fibrin is formed naturally following bone injury. The significance of fibrin in bone healing, discussed above, suggests that fibrin is an importance substance necessary for bone regeneration. Fibrin for instance could be used along with other synthetic materials to fill a bone defect creating an osteoconductive and osteoinductive environment. For example, fibrin mixed with calcium phosphate and applied to critical size defects in the femoral condyle of rabbits showed osteoinductive and osteoconductive properties (Nihouannen et al., 2007). Furthermore, an injectable form of fibrin-osteoblast composite showed enhanced bone regeneration (Kneser et al., 2005). While fibrin does not offer the same mechanical properties as bone, it definitely offers the environment for healthy tissue to grow; and its usage in bone engineering will continue to grow.

20.7.4 In skin Fibrin is used in delivering skin cells to treat chronic wounds. Horch has done tremendous work showing the successful use of fibrin to deliver autologous human keratinocytes to treat chronic wounds (Horch et al., 2001; Vanscheidt et al., 2007). Companies such as Intercytex use fibrin to deliver autologous fibroblasts to treat chronic wounds. A recent publication reported the use of fibrin as a delivery vehicle of autologous mesenchynal stem cells to treat chronic wound (Falanga et al., 2007). Unlike the use of fibrin in bone where fibrin is a component of the construct, in skin, fibrin could be the only structural component needed to deliver cells.

20.8

Conclusions

This review illustrates the versatility of fibrin to deliver growth factors and cells to treat various soft and hard tissue diseases. It is a moldable and easily

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manipulated material to construct biomatrices for tissue engineering, drug and cell delivery, and a host of other biomedical applications. It is a natural scaffold that degrades with minimum rejection, and its degradation can be controlled by modifying its compositions, which makes it an appealing material for many applications. Furthermore, by manipulating the composition of the fibrin components or by addition of other compounds, we can modulate its mechanical and biological properties to custom create products. Thus, one can dial-in the proteins of interest to be secreted by the cells by simply changing the concentration of fibrinogen and thrombin, change its porosity to control rate of growth factors or drug delivery, change its fibril density to inhibit cell migration or block cell invasion, and mix with other natural or synthetic materials to create custom design scaffolds with different mechanical and biochemical characteristics. Clearly, the potential for its application is boundless. The only limitation to its usage is the imagination.

20.9

References

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MacPhee M J, Singh M P, Brady R, Jr. Akhyani N, Liau G, Lasa Jr. C, Hue C, Best A and Drohan W (1996), Fibrin sealant: a versatile delivery vehicle for drugs and biologics, in Surgical Adhesives and Sealants Current Technology and Applications, Sierra D H, Saltz R, eds, Technomic Publishing Co, Lancaster, PA. Mah-Becherel M, Ceraline J, Deplanque G, Bergerat J P, Cazenave J P and Klein-Soyer C (2002), Anti-angiogenic effects of the thienopyridine SR 25989 in vitro and in vivo in murine pulmonary metastasis model, British Journal of Cancer, 86, 803–810. Mana M, Cole M, Cox S and Tawil B (2006), Human U937 monocyte behavior and protein expression on various formulations of three-dimensional fibrin clots, Wound Repair and Regeneration, 14(1), 72–80. Marchac D and Greensmith A L (2005), Early postoperative efficacy of fibrin glue in face lifts: a prospective randomized trial, Plast Reconstr Surg Mar, 115(3), 911–916. Marticorena J, Rodriguez-Ares M T, Tourino R, Mera P, Valladares M J, Martinez-de-laCasa J M and Benitez-del-Castillo J M (2006), Pterygium surgery: conjunctival autograft using a fibrin adhesive, Cornea, 25(1), 34–36. Massone P P B, Magnani B, Conti B, Lequaglie C and Cataldo I (2003), Cauterization versus fibrin glue for aerostasis in precision resections for secondary lung tumors, Annals of Surgical Oncology, 10, 441–446. Matarasso A, Rizk S S and Markowitz J (2005), Short scar face-lift with the use of fibrin sealant, Dermatol Clin, 23(3), 495–504, vii. Meana A, Iglesia J, Del Rio M, Larcher F, Madrigal B, Fresno M F, Martin C, San Roman S and Tevar F (1998), Large surface of cultured human epithelium obtained on a dermal matrix based on live fibroblast-containing fibrin gels, Burns, 24, 621–630. Mesa J M, Zaporojan V, Weinand C, Johnson T S, Bonassar L, Randolph M A, Yaremchuk M J and Butler P E (2006), Tissue Engineering Cartilage with Aged Articular Chondrocytes In Vivo, Plastic & Reconstructive Surgery, 118(1), 41–49. Miller A G, Bottoms S, Laurent G, Carmeliet P and Herrick S (2005), Fibrin-induced skin fibrosis in mice deficient in tissue plasminogen activator, American Journal of Pathology, 167(3), 721–732. Mogford J, Tawil B and Mustoe T (2004), Fibrin sealant combined with fibroblasts and PDGF enhance wound healing in excisional wounds, Wound Repair and Regeneration, 12(2), A9–A9(1). Mosesson M W, Siebenlist K R and Meh D A (2001), The structure and biological features of fibrinogen and fibrin, Ann N Y Acad Sci, 936, 11. Nicosia R F, Tchao R and Leighton J (1983), Angiogenesis-dependent tumor spread in reinforced fibrin clot culture, Cancer Research, 43, 2159–2166. Nihouannen D, Saffarzadeh A, Aguado E, Goyenvalle E, Gauthier O, Moreau F, Pilet P, Spaethe R, Daculsi G and Layrolle P (2007), Osteogenic properties of calcium phosphate ceramics and fibrin glue based composites, J of Materials Science, 18(2), 225–235. Park S H, Park S R, Chung S, Pai K S and Min B H (2004), Tissue-engineered cartilage using fibrin/hyaloran composite gel and its in vivo implantation, Artificial Organ, 29(10), 838–860. Perez R L and Roman J (1995), Fibrin enhances the expression of IL-1 beta by human peripheral blood mononuclear cells, Implications in pulmonary inflammation, J Immunol, 154, 1879–1887. Perka C, Arnold U, Spitzer R S and Lindenhayn K (2001), The use of fibrin beads for tissue engineering and subsequential transplantation, Tissue Eng, 7(3), 359–361. Radosevich M, Goubran H A and Burnouf T (1997), Fibrin sealant: scientific rationale, production methods, properties, and current clinical use, Vox Sang, 72, 133–143.

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Roberts W W, Lorand L and Mockros L F (1973), Viscoelastic properties of fibrin clots, Biorheology, 10, 29. Sade B, Mohr G and Frenkiel S (2006), Management of intra-operative cerebrospinal fluid leak in transnasal transsphenoidal pituitary microsurgery: use of post-operative lumbar drain and sellar reconstruction without fat packing, Acta Neurochir (Wien), 148(1), 13–18, discussion 18–19. Saed G M, Kruger M, Diamond M P (2004), Expression of transforming growth factorβ and extracellular matrix by human peritoneal mesothelial cells and by fibroblasts from normal peritoneum and adhesions: Effect of Tisseel, Wound Repair and Regeneration, 12(5), 557–564. Sahni A and Francis C W (2004), Vascular endothelial growth factor binds to fibrinogen and fibrin and stimulates endothelial cell proliferation, Blood, 96, 3772–3778. Salvino C K, Esposito T J, Smith D K, Jacobs H K, Candel A G, Dries D and Gamelli P (1993), Laparoscopic injection of fibrin glue to arrest intraparenchymal abdominal hemorrhage: An experimental study, J Trauma, 35, 762–766. Scheberich A and Beretz A (2000), Culture of vascular cells in tri-dimensional (3d) collagen: a methodological review, Therapie, 55(1), 35–41. Schense J C, Bloch J, Aebischer P and Hubbell J A (2000), Enzymatic incorporation of bioactive peptides into fibrin matrices enhances neurite extension, Nature Biotechnology, 18, 415–419. Sese N, Cole M, Cox S and Tawil B (2003), Delivering human keratinocytes using fibrin sealant: in vitro studies. The Wound Healing Society, 13th Annual Conference, Seattle, Washington. Sese N, Cole M and Tawil B (2004), Cross-talk between fibroblasts and keratinocytes in 3-D fibrin clots, in vitro Studies, The Wound Healing Society 14th Annual Conference, Atlanta, GA. Sierra D H (1993), Fibrin sealant adhesive systems: A review of their chemistry, material properties, and clinical applications, J Biomater Appl, 7, 309. Silecchi G, Boru C, Moulel J, Rossi M, Anselmino M, Tacchino R, Foco M, Gaspari A, Gentileschi P, Morino M, Toppino M, Basso N (2006), Clinical evaluation of fibrin glue in the prevention of anastomotic leak and internal hernia after laparoscopic gastric by pass: Preliminary results of a prospective, randomized multicenter trial, Obesity Surgery, 16(2), 125–131. Silverman Ronald P, Passaretti David, Huang, Wynne, Randolph, Mark A, Yaremchuk, Michael J (1999), Injectable Tissue-Engineered Cartilage Using a Fibrin Glue Polymer, Plastic & Reconstructive Surgery, 103(7), 1809–1818. Simon D I, Ezratty A M, Francis S A, Rennke H and Loscalzo J (1993), Fibrin(ogen) is internalized and degraded by activated human monocytoid cells via Mac-1 (CD11b/ CD18): a nonplasmin fibrinolytic pathway, Blood, 82, 2414–2422. Sohn D, Lottman L, Lum L, Kim S, Pedowitz R, Coutts R and Sah R L (2002), Effect of Gravity on Localization of Chondrocytes Implanted in Cartilage Defects, Clinical Orthopaedics & Related Research, 394, 254–262. Sweeny S M, Guy C A, Fields G B and San Antonio J D (1998), Defining the domains of type I collagen involved in heparin binding and endothelial tube formation, Pro Natl Acad Sci USA, 95, 7275–7280. Sweeny S M, DiLullo G, Slater S J, Martinez J, Lozzo R V, Lauer-Fields J L, Fields G B and San Antonio J D (2003), Angiogenesis in collagen I requires alpha2-beta1 ligation of a GFP*GER sequence and possibly p38 MAPK activation and focal adhesion disassembly, Journal of Biological Chemistry, 278(33), 30516–30524.

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Tawil B (2005), Fibrin and its applications, Review article. In An Introduction to Biomaterials, Guelcher S A and Hollinger J O (eds), Chapter 7, 105–120 Boca Raton, FL, CRC Taylor & Francis. Tuan T L, Song A, Chang S, Younai S and Nimni M E (1996), In vitro fibroplasia: matrix contraction, cell growth, and collagen production of fibroblasts cultured in fibrin gels, Experimental Cell Research, 223, 127–134. Tuan T L, Zhu J Y, Sun B, Nichter L S, Nimni M E and Laug W E (1996), Elevated levels of plasminogen activator inhibitor-1 may account for the altered fibrinolysis by keloid fibroblasts, J Invest Dermatol, 106, 1007–1011. Vailhe B, Leconte M, Wiernsperger N and Tranqui L (1998), The formation of tubular structures by endothelial cells is under the control of fibrinolysis and mechanical factors, Angiogenesis, 2(4), 331–344. Van Griensven M, Zeichen J, Pape H C, Lehmann U, Bosch U and Seekamp A A (2000), Modified method to culture human osteoblasts using Tissuecol® to improve bone sample adhesion and cellular outgrowth, Cells Tissues Organs, 166, 72–76. Vanscheidt W, Ukat A, Horak V, Bruning H, Hunyadi J, Pavlicek R, Emter M, Hartmann A, Bende J, Zwingers Th, Ermuth T and Eberhardt R (2007), Treatment of recalcitrant venous leg ulcers with autologous keratinocytes in fibrin sealant: A multinational randomized controlled clinical trial, Wound Repair and Regeneration, 15, 308–315. Wechselberger G, Schoeller T, Stenzl A, Ninkovic M, Lille S and Russell R C (1998), Fibrin glue as a delivery vehicle for autologous urothelial cell transplantation onto a prefabricated pouch, The Journal of Urology, 160, 583–586. Werner S and Grose R (2003), Regulation of wound healing by growth factors and cytokines, Physiol Rev, 83, 835–870. Wong C, Inman E, Spaethe R and Helgerson S (2003), Fibrin-based biomaterials to deliver human growth factors, Thromb Haemos, 89, 573–582. Yang Y I, Seol D L, Kim H I, Cho M H and Lee S J (2007), Composite fibrin and collagen scaffold to enhance tissue regeneration and angiogenesis, Current Applied Physics, 7(1), e103–e107. Ye O, Zünd G, Benedikt P, Jockenhoevel S, Hoerstrup S P, Sakyama S, Hubbell J A and Turina M (2000), Fibrin gel as a three dimensional matrix in cardiovascular tissue engineering, Eur J Cardiothorac Surg, 17, 587–591. Young H, Wu B and Tawil B (2007), Mechanical Characteristics of Fibrin and Collagen Scaffolds Impeded with Cells, TERMIS, Toronto, Canada. Yucel E A, Oral O, Olgac V and Oral C K (2003), Effects of fibrin glue on wound healing in oral cavity, J Dent, 31(8), 569–575. Zhang G, Wang X, Wang Z, Zhang J and Suggs L (2006), A PEGylated fibrin patch for mesenchymal stem cell delivery, Tissue Engineering, 12(1), 9–19.

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21 Natural-based polymers for encapsulation of living cells: Fundamentals, applications and challenges P. D E V O S, University Hospital of Groningen, The Netherlands

21.1

Introduction

The introduction of the concept of encapsulation of living cells for treatment of diseases dates back to 1933. The concept was published by Bisceglie (1933) who implanted insulin-producing tissue encapsulated in an immunoprotective membrane in mammals. Unfortunately, Bisceglie (1933) did not recognize the principle applicability of the approach for treatment of disease. He did his experiments to study the effects of the absence of vascularization on the survival of tissues. It took until 1943 before Algire (1943) recognized that graft rejection could be delayed by encapsulating allo- and xenogenic tissues before transplantation. The technology of encapsulation of living cells revisited two decades ago has since grown to a mature research field and is under study for the treatment of a wide variety of diseases, including Hemophilia B (Liu et al., 1993), anemia (Koo and Chang, 1993), dwarfism (Chang et al., 1993), kidney (Cieslinski and Humes, 1994) and liver failure (Uludag and Sefton 1993), pituitary (Colton, 1995) and central nervous system insufficiencies (Aebischer et al., 1994), and diabetes mellitus (Lim and Sun, 1980). In the present review we will discuss the progress of the past two decades. We will start with discussing the different approaches of encapsulation. We will discuss biocompatibility issues in the application of polymers for cell encapsulation since optimal biocompatibility has been a major challenge in the application of cell encapsulation. The different types of molecules applied in encapsulation will be discussed. To give a complete overview the molecules discussed will be of both synthetic and natural origin. The majority of groups nowadays focus on alginate as the principle polymer of choice for fabrication of capsules. Therefore many of the issues in the following section will be discussed in view of the application of alginate in various compositions and forms. Since the vast majority of studies have been performed with pancreatic islets and the cell source of interest, we have mainly focused on this field of application in the present review. 549 © 2008, Woodhead Publishing Limited

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21.2

Approaches of encapsulation: Materials and biocompatibility issues

Encapsulation can be divided into two major designs, viz. intravascular devices and extravascular devices (Figure 21.1). Also there are categories of geometry: tissue can be enveloped in macrocapsules and in microcapsules. The macrocapsules contain groups of cells enveloped together in one immunoisolating membrane that can be implanted as an extravascular and intravascular device. With microencapsulation, the groups of cells or tissues such as pancreatic islets are individually enveloped by their own capsule. These different approaches will be discussed separately in the following sections.

21.2.1 The intravascular approach Intravascular devices have been applied only as macrocapsules in which groups of cells have been enclosed in semipermeable membranes. A major Glucose

Insulin

Intravascular

Extravascular

Macrocapsule

Solid support Blood vessel

Microcapsule

Nutrients

21.1 Immunoisolating devices. In the intravascular macrocapsule, islets are enclosed in a housing surrounding a semipermeable membrane. The device is implanted as a vascular shunt. In the extravascular capsules, islets are enclosed in macrocapsules or enveloped in microcapsules and implanted, without direct vascular connection.

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advantage of intravascular devices is the close contact with the blood stream. As a consequence the islets are in close contact with the blood stream which implies a fast exchange of glucose and insulin. The intravascular device is usually composed of a semipermeable material with blood flow through its lumen. The outside contains the implanted tissue (Chick et al., 1975; Knazek et al., 1972). The device is implanted by vascular anastomoses to the systemic circulation of the host. The most intensively studied intravascular device is the modified diffusion chamber of Chick et al. (1975). It is technically advanced and has been tested extensively in small (Sun et al., 1977) as well as in large animals (Maki et al. 1993; Sun et al., 1977). The original device was composed of a number of small diameter artificial capillaries contained by one large diameter tube. The artificial capillaries were composed of fibers of polyacrylonitrile and polyvinylchloride copolymer (PAN-PVC) similar to those used in extravascular devices (Colton and Avgoustiniatos, 1991; Lanza et al., 1992). This PAN-PVC ultrafiltration capillary design (Colton, 1995) has a lumen loaded with pancreatic islets between the outside of the artificial capillaries. The design permits close contact between the islets and blood, separated only by the microporous walls of the capillaries. These devices were found to induce normoglycemia in diabetic rats (Sun et al., 1977), dogs (Maki et al., 1993), and monkeys (Sun et al., 1977) but required systemic anticoagulation. The duration of this normoglycemia was usually restricted to several hours and successes of a somewhat longer duration were exceptional. Clotting of the blood in the lumen of these small diameter artificial capillaries proved to be a major obstacle, in spite of anticoagulant medication in massive doses. This thrombus formation was an early sign of insufficient biocompatibility, and has led to the use of tubular membranes with larger diameters in the hope of minimizing or eliminating clot formation in the absence of systemic anticoagulation. The latter large lumen device is composed of a single, coiled, and tubular membrane with an internal diameter of 5–6 mm. The membrane is somewhat modified but still composed of PAN-PVC with a nominal molecular weight cutoff of 50 kDa. This membrane was found to be rather successful, since these devices implanted as high flow arteriovenous fistulas could remain patent for periods of seven weeks in the absence of systemic anticoagulant therapy (Galletti et al., 1981). This success is in part explained by the high flow rates through the device which prevents adhesion of cells to the membranes or collection of those cells in the immediate vicinity (Colton and Avgoustiniatos 1991). Allo- and xenogenic islets in the high flow devices were successfully transplanted to diabetic dogs (Lanza et al., 1993; Lanza and Chick, 1997a; Lanza and Chick, 1997b; Maki et al., 1995; Maki et al., 1996a; Maki et al., 1996b) but the efforts to improve the blood-compatibility have probably interfered with the efficacy of the device as an implantation site for islets.

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This view is derived from the following observations. First, two devices per recipient instead of one were required to achieve adequate secretion capacity while maintaining the same numbers of islets per device (Lanza et al., 1992). Furthermore, it has not been possible to load the space between the membranes and the housing with an islet-tissue density higher than 5– 10% of the volume (Colton, 1995), in spite of the fact that the large lumen is exposed to arterial blood with optimal concentrations of nutrients and oxygen. It is quite plausible that the high flow rates through the device, which are required to keep the device patent, do not allow sufficient exchange of glucose, insulin and nutrients to permit long term survival and adequate function of the islets. There are also indications that the materials applied in this kind of device are not only thrombogenic but also insufficiently compatible with long term functional survival of the cells and tissues. For example, the PTFE as used for vascular anastomosis of the device has been shown to induce IL-1β production by macrophages (Krause et al., 1990), which cytokine is lethal for islets (Cetkovic Cvrlje and Eizirik, 1994; Sandler et al., 1994). It is quite plausible that IL-1β causes loss of high numbers of islets during the period between implantation and complete integration of the prosthesis since macrophages are usually the first cells to invade the implant (Anderson and Langone 1999; Hunt et al., 1996; Remes and Williams 1992). This is another explanation for the fact that so many islets divided over two devices are required for maintaining normoglycemia in dogs. Although the intravascular devices have shown some degree of success, the problems mentioned above should be solved if clinical application is considered. Even then, the complications associated with any type of vascular prosthetic surgery remain a serious threat, such as thrombosis, either primary or secondary to intimal hyperplasia at the venous anastomosis, defects of the device, or infection. This is a major drawback for wide application in large numbers of diabetic patients since any alternative to conventional insulin treatment should preferably carry no additional risk.

21.2.2 The extravascular approach Due to the potential risks of connecting a device to the host blood stream many have abandoned the intravascular systems and moved to the extravascular approach during the past decade. The concept of extravascular devices does not require vascular anastomoses, since it is based on the principle of diffusion chambers (Scharp et al., 1984). The extravascular devices can be implanted with minimal surgery in different sites and are not associated with major risks such as thromboses. Extravascular devices can be categorized into two different types of devices, viz. the extravascular macrocapsules and extravascular microcapsules.

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The extravascular macrocapsules can be implanted with minimal surgery in different sites such as the peritoneal cavity (Archer et al., 1980; Jain et al., 1999; Lanza et al., 1994; Loudovaris et al., 1999), the subcutaneous site (Juang et al., 1994; Juang et al., 1995; Lacy et al., 1991; Scharp et al., 1994; Suzuki et al., 1998a; Suzuki et al., 1998b; Tatarkiewicz et al., 1999), or the renal capsule (Siebers et al., 1990). The geometry of macrocapsules may be planar in the form of a flat, circular double layer or tube-like as a so-called hollow fiber (Scharp et al., 1984). The extravascular macrocapsules can be readily retrieved. Due to the advantages over other systems, macrocapsules have been intensively studied. Different geometries have been tested. The tube geometry is usually preferred over the planar membranes for its higher degree of biocompatibility (Woodward, 1982). Most studies on tube-shaped fibers use fibers of polyacrylonitrile and polyvinylchloride copolymer (PAN-PVC) (Colton and Avgoustiniatos, 1991; Lanza et al., 1993). They have been produced with a smooth or fenestrated outer layer. The design with the smooth outer skin proved to be the most successful since it provokes much less fibrosis than the rough fenestrated surface which allows host tissue to grow into the spongy matrix. Many modifications of this concept have been proposed in order to further improve the biocompatibility. One of those was the coating of the membranes with poly-ethylene-oxide to reduce protein adsorption (Shoichet et al., 1994). Biocompatibility problems with extravascular macrocapsules are usually deleterious only to the function of the encapsulated tissue and have no or only minimal risk for the recipient. These biocompatibility problems are usually related to toxicity and activation of non-specific foreign body reactions resulting in fibrotic overgrowth with subsequent necrosis of the encapsulated tissue. In the past decade many groups have studied the applicability of hydrogels for extravascular macroencapsulation. Hydrogels provide a number of features which are advantageous for the biocompatibility of the membranes. First, as a consequence of the hydrophilic nature of the material, there is almost no interfacial tension with surrounding fluids and tissues which minimizes the protein adsorption and cell adhesion. Furthermore, the soft and pliable features of the gel reduce the mechanical or frictional irritations to surrounding tissue (De Vos and Tatarkiewicz, 2002; De Vos and Van Schilfgaarde, 1999). The materials applied for the hydrogels are polyamide (Isayeva et al., 2003; Lhommeau et al., 1997), alginate (Lanza et al., 1996; Lanza and Chick 1995; Lanza et al., 1996), agarose (Iwata et al., 1992; Jain et al., 1999), 2hydroxyethyl methacrylate HEMA (Klomp et al., 1979; Sefton 1993), and a copolymer of acrylonitrile and sodium-methallyl sulfonate, AN69 (Kessler et al., 1991). Some supporting results have been shown with the hydrogel membrane AN69, which induced only minimal fibrosis in the peritoneal cavity of rats (Kessler et al., 1995; Kessler et al., 1997). Kessler et al. (1995, 1997)

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have introduced corona discharge to obtain a membrane with a more hydrophobic surface in order to facilitate diffusion of insulin. Fewer molecules adhered to the surface of such membranes, which improved not only the permeability for insulin but also its long term biocompatibility. One year after implantation of empty capsules in rats, only few macrophages were found on the membrane’s surface. Surprisingly, up to now, there are no in vivo data available on the efficacy of the membranes in facilitating survival of encapsulated pancreatic islets. It seems that many groups have abandoned AN69 and have focused their research efforts on membranes prepared from polyvinylalcohol (PVA) which have been shown to allow for long-term survival of islet-tissue (Burczak et al., 1996; Qi et al., 2004). Recently, it has been reported that a commercial available macrocapsule, viz. TheraCyt, allows for survival of pancreatic islets in cynomolgus monkeys for period up to 8 weeks (Elliott et al., 2005). Also, unconfirmed reports mention successful treatment of type I diabetic patients with macroencapsulated porcine islets. A 3-cm stainless-steel mesh capsule containing a removable Teflon cylinder was inserted into the abdominal cavity of each patient. Two months later, after a collagen membrane had formed around the capsule, the cylinder was removed and approximately one million pig islets were injected into the tube. The mixture of cells consisted of islet cells and testicular Sertoli cells taken from neonatal pigs. Sertoli cells are thought to have a special ability to subdue immune system T cells, which normally fight against the presence of anything foreign. The researchers gave the cells to 12 children with type 1 diabetes between the ages of 11 and 17 and did not administer any immunosuppressive drugs to protect the cells from being rejected. Six of the 12 patients have functioning grafts, and subsequent to receiving additional transplanted islets at 20 weeks, they have demonstrated improvements in their islet function. According to the researchers, the pig cells did not elicit the expected immune system response in the patients, and retransplantation of islets failed to stimulate a secondary response against the pig cells as well. One child remains insulin-independent one year after the transplant. Another was insulin-independent for six months and now requires 75% less insulin than before the procedure. A novel, innovative approach in the field of macroencapsulation is the introduction of so-called ‘smart’ membranes (Isayeva et al., 2003; Kurian et al. 2003). These smart-membranes are composed of tricomponent amphiphilic membranes containing poly(ethylene glycol) (PEG), polydimethylsiloxane (PDMS) and polypentamethylcyclopentasiloxane (PD(5)). It has been shown that these membranes in air are enriched by the hydrophobic components, PDMS and PD(5), while in water the surfaces are rich in the hydrophilic PEG. This illustrates the versatile properties of the membranes under different circumstances. It has been shown that these versatile properties facilitate diffusion of essential nutrients such as oxygen and of therapeutic hormones

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such as insulin. Also it has been shown that the membranes allow for long term survival of encapsulated islet tissue. All the above mentioned studies demonstrate the principle clinical applicability of the technology of extravascular macro-encapsulation technology. However, the same studies also show that the technique is not yet ready for clinical application since graft survival is usually limited to several months. The vast majority of research groups nowadays focus on microcapsules instead of on macrocapsules. The reason is the theoretical advantages of microcapsules. The spherical shape of microcapsules offers better diffusion capacity because of a better surface/volume ratio. Microcapsules cannot be easily disrupted, they are mechanically stable, they do not require complex or expensive manufacturing procedures, and they can be implanted into the patient by a simple injection procedure. Due to the flexible and pliable characteristics, microcapsules are mostly produced from hydrogels. Most of the work has been done with alginatebased microcapsules, a natural polymer. As a consequence of the large experience with alginate, the vast majority of the research discussed below will be dedicated to alginate-based microcapsules. Microcapsules have been shown not to interfere with cellular function and alginate-based capsules have been shown to be stable for years in small and large animals and also in men (Soon Shiong et al., 1994). The technique is based on entrapment of individual islets in a droplet which is transformed into a rigid capsule by gelification and/or subsequent coating to provide semipermeable properties. During the past decade our group has concentrated mainly on the alginatepoly-l-lysine (PLL) system. An important aspect for preparing successful alginate-PLL capsules, which requires a lot of experience in the microencapsulation process, is a smooth and mechanically stable microcapsule. Alginates can be obtained in different compositions. These alginate vary in their guluronic-acid (G) and mannuronicacid content We have observed that high rather than low viscosity alginates produce smooth beads with no obvious tails or strains. Moreover, we have found that after implantation, alginates with low-G concentrations had a tendency to swell with subsequent breakage of the PLL membrane followed by cellular overgrowth of the capsules. Therefore, for our studies we have chosen the alginate with an intermediate G-concentration. The stability of capsules is determined by the stability of a membrane and the stability of an alginate core. The stability of a membrane can be controlled by the PLL-step since shorter incubation time, lower PLL concentrations, and lower PLL molecular weight than described above, were associated with an increase of the number of capsules with broken membranes (Bunger et al., 2005; De Haan et al., 2003; De Vos et al., 2003a; De Vos et al., 1996a; De Vos et al., 1996b; De Vos and Van Schilfgaarde, 1999). Additionally, incubation at 4°C

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instead of room temperature led to less stable microcapsules. Finally, we have not applied EGTA or citrate (Fritschy et al., 1991) to liquefy the inner core of the capsule. The reason for this modification of the original Lim and Sun method (Lim and Sun, 1980) was that we have observed many capsules losing their integrity during the treatment. After the procedure the capsules should have an adequate rigidity. They should be able to withstand the shearforces they can experience in vivo. Durability tests are rather conflicting in their predictions on the required durability; it is in the range of 2–10 g per capsule (Lacík 2006). For many years researchers have considered insufficient biocompatibility to be a major threat for clinical application of microcapsules. Failure of microencapsulated islet grafts was usually interpreted to be the consequence of insufficient biocompatibility of the materials applied, which induces a non-specific foreign body reaction against the microcapsules and results in progressive fibrotic overgrowth of the capsules. This overgrowth interferes with adequate nutrition of the islets and consequently causes islet cell death. During the past decade many improvement have been introduced in encapsulation technology which have solved the vast majority of bioincompatibility problems. It was shown that pure alginate rather than commercially available crude alginates should be applied for encapsulation (De Vos and Tatarkiewicz, 2002). Crude alginate was shown to be associated with overgrowth of the capsules by inflammatory cells (mostly macrophages and fibroblasts), with necrosis of the enveloped therapeutic cells as a consequence.

21.3

Physico-chemistry of microcapsules and biocompatibility

An analysis of the type of studies on encapsulation learns that until 2000 most groups were investigating the functional performance of immunoisolated grafts in order to demonstrate ‘the proof of principle’. Usually these groups demonstrated survival for prolonged but limited time periods of allogenic and xenogenic grafts in the absence of immunosuppression. None of these groups (including ours) were able to explain the limitations in survival time. From 2000 a number of groups changed their strategy and started basic studies toward the factors determining success and failure of encapsulated grafts. As a consequence new physicochemical technologies have come to the field during the past five years. This has brought novel insights. In order to provide more insight into the structure of alginate-PLL capsules we have performed a physico-chemical analysis of the capsules by applying X-ray photoelectron spectroscopy (De Vos et al., 2002a; De Vos et al., 2002b; Van Hoogmoed et al., 2003). This technique allows for identification of the chemical groups on the surface of the capsule on an atomic level. Up to a

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few years ago the capsule was assumed to be composed of a core of Caalginate which is enveloped by a membrane composed of two layers, viz. an inner layer of alginate-PLL and an outer layer of Ca-alginate (Dupuy et al., 1994; Lim and Sun, 1980; Thu et al., 1996). The data which have led to this model were almost exclusively obtained by studying the chemical interactions of PLL by rinsing non-bound Ca2+ and often individual components of alginate (i.e guluronic (G)-acid and mannuronic (M)-acid monomers) and not by studying the chemical structure of the capsules as such. In our studies on true capsules, we never observed Ca2+ in the membrane of the capsules, which has the following implications for the assumed structure of the capsules. First, the alginate-PLL layer is not composed of the combination of alginatePLL and Ca2+-alginate but of the alginate-PLL only. The absence of Ca2+ also implies that the outer Ca-alginate layer does not exist and, consequently, that the membrane is not composed of two but of one layer only. Finally, we found sodium in the membrane, which is bound by carboxyl groups on consecutive blocks of G- and M-molecules, which remain unbound between the complexes of PLL and mixed sequences of G and M. Figure 21.2 shows the actual structure of alginate-PLL capsules. Since then many new physico-chemical technologies have come to the field which has been very helpful in understanding and explaining biocompatibility issues in encapsulation. These techniques are Fourier transform infrared spectroscopy (FTIR) (Van Hoogmoed et al., 2003), X-ray photoelectron spectroscopy (XPS) (De Vos et al. 2003b; Ponce et al., 2006), with time-offlight secondary ion mass spectrometry (ToF-SIMS) (Tam et al., 2005), atomic force microscopy (Bunger et al., 2003b), and recently zeta-potential measurements (De Vos et al., 2007). These new technologies have changed the way of thinking in the field. To demonstrate this we will discuss the structural features of the capsules in somewhat more depth. The finding that capsules are not having an outer alginate layer implies that the proinflammatory PLL is always on the surface of the capsules in direct contact with the inflammatory cells. This suggests that, for optimal biocompatibility, we have to focus on understanding and improving the interaction of the inflammatory polycations with alginate rather than improving the second coating step with alginate. Alternatively, we can perform an additional envelopment step of the capsules to prevent direct contact between the polycation with the surrounding tissues in the implantation site (Bunger et al., 2003a). Physicochemical differences in capsule types are not always measurable on fresh capsules. For many years, we and others have been trying to find a technology to measure charge density differences on alginate-PLL capsules composed of intermediate-G and high-G alginates. We were not able to demonstrate a difference with fresh capsules. We required capsules that had been implanted for a few days to demonstrate such a difference. At that moment we realized that the in vivo circumstances may change the capsule

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Outer Ca-alginate layer A. 2-layer membrane model

M–M–M–M COOH O O HO HO

Ca-alginate core

HO O

HO COOH O

O

COOH O HO HO

HO O

HO COOH

G–G–G–G O

HO

COOH OH OH O

O

COOH HO

O O OH C HO O O H

OH O

O

O O OH C O O H

B. The actual membrane structure

21.2 The considered and the actual structure of alginate-PLL capsules. The capsule is not composed of three layers as generally assumed but of two layers. © 2008, Woodhead Publishing Limited

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Inner alginate-PLL layer

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properties such that the capsule become more prone to inflammatory responses. Up to now it has been insufficiently realized that the direct environment of the capsules changes directly after implantation. A pertinent change is a drop in pH as the consequence of a temporary inflammation process due to the mandatory surgery. Such a drop in pH can for instance induce changes in the charge density of the capsules and make the capsule more vulnerable for adhesion of proteins and cells. Capsules should be able to withstand these kinds of environmental changes. We found that high-G capsules showed statistically significant more positive charges at lower pH than intermediateG capsules, which corresponds with the higher degree of biocompatibility of intermediate-G capsules (De Vos et al., 2006). In a recent study, we applied all the current knowledge for the requirements of producing a biocompatible alginate-PLL capsule. These capsules were implanted in the peritoneal cavity of rats and retrieved two years later, i.e. the life-span of a rat, for studying the biocompatibility of the capsules. It was found that the vast majority of the capsules could be retrieved after this twoyear period. Of the retrieved capsules only a portion of 2–10% was overgrown with inflammatory cells while 90–98% of the alginate-PLL capsules were completely free of any inflammatory overgrowth (De Vos et al. 2002a). This study shows that it is feasible to produce fully biocompatible alginate-PLL capsules in spite of the inflammatory reactions individual components of the capsules can provoke.

21.4

Immunological considerations

During recent years it has become recognized that immunological responses against immunisolating devices is far more complicated than only a response against the capsule materials. The immunological responses can be directed against the capsule but also against the cells in the capsules. Up to now at least four different responses have been identified. The first is a nonspecific activation of the innate immune system by the surgical procedure of transplantation. The second is the foreign body response against the capsule. The third type of response is provoked by the enveloped tissue which releases bioactive factors but also allogenic or xenogenic epitopes. The reaction of the host immune system towards the capsule and the encapsulated tissue is both through the innate and the adaptive lineages. The last identified type of response is the deleterious component of the vascularization process which only applies for capsule types which will be vascularized after implantation. The activation of the innate immune system already starts with the mandatory surgery to implant the encapsulated cells. This mandatory surgery induces an inflammatory response due to rupture of bloodvessels which is associated with influx of inflammatory cells and release of bioactive factors such as cytokines and fibronectin. It depends on the material’s properties whether

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this results in adsorption of proteins and subsequently cell adherences onto the surface. The second response, the foreign body response against the capsules, can now start, depending on the characteristics of the materials applied. Fibrosis which affects the whole graft and not only a portion of the capsules has become a rare phenomenon since the introduction of purified alginates. Recently, a detailed study on the tissue responses against alginate-PLL capsules has been published by Robitaille et al. (2005). The authors observed a strong fibrogenic response with high concentrations of fibrogenic cytokines such as TGF-Beta, in contrary to our results It should be noted, however, that the alginates applied had a low purity degree and therefore a low biocompatibility. There is obviously a large difference in pathophysiology of the reaction between application of capsules with a high and a low degree of biocompatibility since we never observe these responses. When applying capsule grafts in which the majority of the capsules remain free of overgrowth, we observe a temporary increase in macrophages, granulocytes and cytokines characteristic for an activation of the innate immunity (Bunger et al. 2003b; De Vos et al., 2002b). This response is usually extinguished within two weeks but, unfortunately, is responsible for loss of a significant number of the encapsulated cells (De Vos et al., 1999; De Vos et al. 2002b). Immune cells such as circulating and tissue specific macrophages and granulocytes can take up components of the foreign material or specific allogenic and/or xenogenic epitopes and initiate a specific immune response characterized by the presence of lymphocytes in the vicinity of the materials. It has been shown that this induces the formation of encapsulated tissue specific antibodies (Lanza et al., 1994; Lanza et al., 1995; Lanza and Chick 1997a) . Most groups however do not considered the formation of antibodies to be deleterious for the tissue since the capsules should adequately protect the tissue against complement and antibody mediated destruction. Release of immune stimulating factors by the encapsulated tissue has recently been identified as a causal factor for 60% of the loss of tissue in the first months after transplantation (De Vos et al., 1999). A conceivable approach to overcome the problem of islet-derived cytokines is reduction of the capsule permeability. The diffusion of graft-derived and inflammatory cell-derived cytokines could be a major threat for the longevity of the encapsulated grafts (Bolzan and Bianchi, 2002; De Vos et al., 2003a). The permeability of the capsules for cytokines has always been a subject of debate. Sceptics have always assumed that the membranes of capsules cannot adequately protect against deleterious cytokines with an approximate molecular weight of insulin or essential nutrients (5–15 kDa). Therefore, diffusion of cytokines into the capsules has always been the Achilles heel of encapsulation. Combined efforts of the De Vos-group and that of Marchetti in Pisa have shown that this is not an insurmountable problem (De Vos and Marchetti 2002). It has been shown that the final effect of cytokines is dependent on the combined

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presence of different cytokines and on the concentration of cytokines (De Vos and Marchetti, 2002). It was found in vitro that decreasing the permeability by chemical modification of the capsules prevents diffusion into the capsules of large and multimeric cytokines such as TNF-α. Also, diffusion of small cytokines (e.g. IL-1β) was reduced by changing the permeability of the membrane, which was unexpected as the molecular weight of small cytokines such as IL-1β (17 kDa) was far below the molecular cut-off of the applied capsules (100 kDa) (Figure 21.3). The mechanism is that the chemical modifications for decreasing the permeability increase the negative charges of the alginate in the capsules which induces repulsion of the negative charges on the tested cytokines-molecules. Also, we found evidence that cytokines may not interfere with islet function in the case of xenografting of encapsulated islets in humans. We have observed that following exposure to a combination of human cytotoxic cytokines, a marked decrease in functional survival and a high percentage of apoptotic cells could be found in human islets but not in bovine islets (Piro et al., 2001). Preliminary data show that this is due, at least in part, to the fact that xenogenic islet cells are less capable to bind and to take up human cytokines. This implies that, in at least in some combinations, even when capsules are applied which are permeable for cytokines, the function and survival of xenogenic islet-sources will be less affected. Some apply specific capsule materials to enforce vascularization of the capsule membranes in order to promote exchange of nutrients and therapeutic agents between the bloodstream and the encapsulated tissue (Cruise et al., 1998; Cruise et al., 1999; Trivedi et al., 2000a; Trivedi et al., 2000b; Yoon et al., 1999). This vascularization of a membrane is preceded by an inflammation episode which involves recruitment of many deleterious inflammatory cells in the vicinity of the capsules and with the formation of an extracellular matrix to facilitate ingrowth of endothelial cells (Auguste et al., 2005; Bonnet and Walsh, 2005; De Vos and Tatarkiewicz, 2002). The latter episode is not only associated with the presence of many deleterious cytokines and bioactive molecules but also with a period of ischemia. It is advisable to apply prevascularized approaches (De Vos et al., 1997; De Vos and Tatarkiewicz, 2002) in order to reduce above mentioned deleterious effects.

21.5

Conclusions and future trends

The concept of ‘smart’ membranes (Isayeva et al., 2003; Kurian et al., 2003) in the field of extravascular macroencapsulation has brought new optimism into the field since the versatile properties of these membranes may facilitate transport of nutrient and other essential molecules to the cells in the relatively large capsules without interfering with immunoprotective properties of macrocapsules. Although the number of scientific publications on these new

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Receptor IL-1β Receptor TNF-α

Bioactive factors MCP-1, 1L-6, NO

Antibodies and immune cells Human IL-1β

Macrophage

Human TNF-α (a) Allogenic

100

Diffusion through capsule (%)

Bioactive factors MCP-1, 1L-6, NO

80 Porosity for secreted insulin 60

40 Porosity for IL-1β 20

0 No capsule

Conventional permeability

Adjusted permeability

(b) Allogenic with adjusted permeability

21.3 Approaches to prevent deleterious effects of diffusion of cytotoxic cytokines and chemokines into the immunoisolating capsules after transplantation. (a) Islets produce cytokines that diffuse out of the capsules and activate inflammatory cells such as macrophages in the vicinity of the microencapsulated islets. The cytokines secreted by the macrophages diffuse into the capsules and induce massive cell death in the allogenic human islet cells. (b) By adjusting the porosity of the capsules for IL-1-β and for secreted insulin, we found that the permeability can be lowered so that cytokines cannot pass the membrane, whereas the porosity of the capsules for insulin remains unaffected (graph show mean ± sem of 5–7 experiments). (c) In the case of xenotransplantation, we found that cytokines of human origin are less deleterious for islets of animal origin such as bovine islets, possibly due to reduced interaction at the receptor level.

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Bioactive factors MCP-1, 1L-6, NO

(c) Xenogenic with reduced capacity of human cytokine to bind to non-human receptor

21.3 (Continued)

materials are still limited, it is predictable that the current attention for nanotechnology will lead to more ‘smart’ membranes and increased insight into the requirements these membranes have to meet in order to allow for adequate survival of the enveloped tissues. It is conceivable that this will lead to a renewed interest in macroencapsulation. Currently, all signs point toward microcapsules as the principle technology for future clinical application of cell-encapsulation. That this is not only the view of the present author is illustrated by the present first phase clinical studies that have recently been started in different places in the world. Calafiore et al. (2006) has started clinical trials in ten diabetic patients in Italy. In his first report he shows the efficacy of the approach in humans since he showed that islets survived for periods up to a year in the first two patients. It should be mentioned though that the patients were still receiving insulin therapy due to insufficient islet mass. In Canada, AmCyte Inc. announced approval for performing clinical trials in 12 patients. Finally, in the US, Novocell, Inc. is in the process of performing clinical trials in 12 patrients (Lacík, 2006). A factor that has contributed to this optimism is the understanding of the basics of microencapsulation and the new insights in requirements capsules have to meet in order to provide an optimal chance of functional survival of the enclosed tissue. The introduction of novel physico-chemical technologies in the field has had, as discussed in the previous sections, a boosting effect on this new knowledge.

21.6

Sources of further information and advice

The European Union has granted a program entitled ‘Bioencapsulation multiscale interaction analysis’. The program involves more than 40 academical and industrial groups working on different aspects of encapsulation. The

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common goal of the groups involved is to improve the knowledge on bioencapsulation in developing reliable, economical and safe industrial encapsulation processes and applications by connecting researchers and industrials in bioencapsulation and promoting exchanges and collaborations. The action is divided into four working groups; it involves separately studies of capsules on: (a) a molecular level; (b) a microcapsules level; (c) a technological level; and (d) an application level. More information and recent progress on the field of cell encapsulation can be obtained from the website: http://cost865.bioencapsulation.net/

21.7

References

Aebischer P, Goddard M, Signore A P and Timpson R L, ‘Functional recovery in hemiparkinsonian primates transplanted with polymer-encapsulated PC12 cells’, Exp Neurol, 1994, 126, 151–158. Algire G H, ‘An adaption of the transparent chamber technique to the mouse’, J Natl Cancer Inst, 1943, 4, 1–11. Anderson J M and Langone J J, ‘Issues and perspectives on the biocompatibility and immunotoxicity evaluation of implanted controlled release systems’, J Controlled Release, 1999, 57(2), 107–113. Archer J, Kaye R and Mutter G, ‘Control of streptozotocin diabetes in Chinese hamsters by cultured mouse islet cells without immunosuppression: a preliminary report’, J Surg Res, 1980, 28, 77–85. Auguste P, Lemiere S, Larrieu-Lahargue F and Bikfalvi A, ‘Molecular mechanisms of tumor vascularization’, Crit Rev Oncol Hematol, 2005, 54(1), 53–61. Bisceglie V V, ‘Uber die antineoplastische Immunitat’, Krebsforsch, 1933, 40, 141–158. Bolzan A D and Bianchi M S, ‘Genotoxicity of streptozotocin’, Mutat Res, 2002, 512(23), 121–134. Bonnet C S and Walsh D A, ‘Osteoarthritis, angiogenesis and inflammation’, Rheumatology (Oxford), 2005, 44(1), 7–16. Bunger C M, Gerlach C, Freier T, Schmitz K P, Pilz M, Werner C, Jonas L, Schareck W, Hopt U T and De Vos P, ‘Biocompatibility and surface structure of chemically modified immunoisolating alginate-PLL capsules’, J Biomed Mater Res, 2003b, 67A(4), 1219– 1227. Bunger C M, Gerlach C, Freier T, Schmitz K P, Pilz M, Werner C, Jonas L, Schareck W, Hopt U T and De Vos P, ‘Biocompatibility and surface structure of chemically modified immunoisolating alginate-PLL capsules’, J Biomed Mater Res, 2003a, 67A(4), 1219– 1227. Bunger C M, Tiefenbach B, Jahnke A, Gerlach C, Freier T, Schmitz K P, Hopt U T, Schareck W, Klar E and De Vos P, ‘Deletion of the tissue response against alginate-pll capsules by temporary release of co-encapsulated steroids’, Biomaterials, 2005, 26(15), 2353–2360. Burczak K, Gamian E and Kochman A, ‘Long-term in vivo performance and biocompatibility of poly(vinyl alcohol) hydrogel macrocapsules for hybrid-type artificial pancreas’, Biomaterials, 1996, 17(24), 2351–2356. Calafiore R, Basta G, Luca G, Lemmi A, Montanucci M P, Calabrese G, Racanicchi L, Mancuso F and Brunetti P, ‘Microencapsulated pancreatic islet allografts into

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nonimmunosuppressed patients with type 1 diabetes: first two cases’, Diabetes Care, 2006, 29(1), 137–138. Cetkovic Cvrlje M and Eizirik D L, ‘TNF-alpha and IFN-gamma potentiate the deleterious effects of IL- 1 beta on mouse pancreatic islets mainly via generation of nitric oxide’, Cytokine, 1994, 6, 399–406. Chang P L, Shen N and Westcott A J, ‘Delivery of recombinant gene products with microencapsulated cells in vivo’, Hum Gene Ther, 1993, 4, 433–440. Chick W L, Like A A and Lauris V, ‘Beta cell culture on synthetic capillaries: an artificial endocrine pancreas’, Science, 1975, 187, 847–849. Chick W L, Like A A, Lauris V, Galletti P M, Richardson P D, Panol G, Mix T W and Colton C K, ‘A hybrid artifical pancreas’, Trans Am Soc Artif Intern Organs, 1975, 21, 8–15. Cieslinski D A and Humes H D, ‘Tissue engineering of a bioartificial kidney’, Biotechnol Bioeng, 1994, 43, 678–681. Colton C K, ‘Implantable biohybrid artificial organs’, Cell Transplantation, 1995, 4, 415–436. Colton C K and Avgoustiniatos E S ‘Bioengineering in development of the hybrid artificial pancreas’, J Biomech Eng, 1991, 113, 152–170. Cruise G M, Hegre O D, Lamberti F V, Hager S R, Hill R, Scharp D S and Hubbell J A, ‘In vitro and in vivo performance of porcine islets encapsulated in interfacially photopolymerized poly(ethylene glycol) diacrylate membranes’, Cell Transplant, 1999, 8(3), 293–306. Cruise G M, Hegre O D, Scharp D S and Hubbell J A, ‘A sensitivity study of the key parameters in the interfacial photopolymerization of poly(ethylene glycol) diacrylate upon porcine islets’, Biotechnol Bioeng, 1998, 57(6), 655–665. De Haan B J, Faas M M and De Vos P, ‘Factors influencing insulin secretion from encapsulated islets’, Cell Transplant, 2003, 12(6), 617–625. De Vos P, De Haan B J, Kamps J A, Faas M M and Kitano T, ‘Zeta-potentials of alginatePLL capsules: A predictive measure for biocompatibility?’, J Biomed Mater Res A, 2007, 80(4), 813–819. De Vos P, De Haan B J, Pater J and Van Schilfgaarde R, ‘Association between capsule diameter, adequacy of encapsulation, and survival of microencapsulated rat islet allografts’, Transplantation, 1996a, 62(7), 893–899. De Vos P, De Haan B J, Wolters G H J and Van Schilfgaarde R, ‘Factors influencing the adequacy of microencapsulation of rat pancreatic islets’, Transplantation, 1996b, 62, 888–893. De Vos P, Hillebrands J L, De Haan B J, Strubbe J H and Van Schilfgaarde R, ‘Efficacy of a prevascularized expanded polytetrafluoroethylene solid support system as a transplantation site for pancreatic islets’, Transplantation, 1997, 63, 824–830. De Vos P and Marchetti P, ‘Encapsulation of pancreatic islets for transplantation in diabetes: the untouchable islets’, Trends Mol Med, 2002, 8(8), 363–366. De Vos P, Smedema I, van Goor H, Moes H, van Zanten J, de Leij L F M and De Haan B J, ‘Association between macrophages activation and function of microencapsulated islets’, Diabetologia, 2003a, 46(5), 666–673. De Vos P and Tatarkiewicz K, ‘Considerations for successful transplantation of encapsulated pancreatic islets’, Diabetologia, 2002, 45, 159–173. De Vos P, Van Hoogmoed C G and Busscher H J, ‘Chemistry and biocompatibility of alginate-PLL capsules for immunoprotection of mammalian cells’, J Biomed Mater Res, 2002a, 60, 252–259.

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De Vos P, Faas M M, Strand B and Calafiore R, Alginate-based microcapsules for immunoisolation of pancreatic islets, Biomaterials, Nov, 27(32), 5603–17. De Vos P, Van Hoogmoed C G, De Haan B J and Busscher H J, ‘Tissue responses against immunoisolating alginate-PLL capsules in the immediate posttransplant period’, J Biomed Mater Res, 2002b, 62(3), 430–437. De Vos P, Van Hoogmoed C G, van Zanten J, Netter S, Strubbe J H and Busscher H J, ‘Long-term biocompatibility, chemistry, and function of microencapsulated pancreatic islets’, Biomaterials, 2003b, 24(2), 305–312. De Vos P and Van Schilfgaarde R, ‘Biocompatibility Issues’, in Cell Encapsulation Technology and Therapeutics, Kühtreiber W M, Lanza R P and Chick W L, eds., Birkhäuser, Boston, 1999, 63–79. De Vos P, van Straaten J F, Nieuwenhuizen A G, de Groot M, Ploeg R J, De Haan B J and Van Schilfgaarde R, ‘Why do microencapsulated islet grafts fail in the absence of fibrotic overgrowth?’, Diabetes, 1999, 48, 1381–1388. Dupuy B, Arien A and Perrot Minnot A, ‘FI-IR of membranes made with alginatepolylysine complexes. Variations with mannuronic or guluronic content of the polysaccharides’, Art Cells, Blood subs, and Immob Biotech, 1994, 22(1), 71–82. Elliott R B, Escobar L, Calafiore R, Basta G, Garkavenko O, Vasconcellos A and Bambra C, ‘Transplantation of micro- and macroencapsulated piglet islets into mice and monkeys’, Transplant Proc, 2005, 37(1), 466–469. Fritschy W M, Wolters G H and Van Schilfgaarde R, ‘Effect of alginate-polylysinealginate microencapsulation on in vitro insulin release from rat pancreatic islets’, Diabetes, 1991, 40, 37–43. Galletti P M, Trudell L A, Panol G, Richardson P D and Whittemore A, ‘Feasibility of small bore AV shunts for hybrid artificial organs in nonheparinized beagle dogs’, Trans Am Soc Artif Intern Organs, 1981, 27, 185–187. Hunt J A, Flanagan B F, McLaughlin P J, Strickland I and Williams D F, ‘Effect of biomaterial surface charge on the inflammatory response: evaluation of cellular infiltration and TNF alpha production’, J Biomed Mater Res, 1996, 31(1), 139–144. Isayeva I S, Kasibhatla B T, Rosenthal K S and Kennedy J P, ‘Characterization and performance of membranes designed for macroencapsulation/implantation of pancreatic islet cells’, Biomaterials, 2003, 24(20), 3483–3491. Iwata H, Takagi T and Amemiya H, ‘Agarose microcapsule applied in islet xenografts (hamster to mouse)’, Transplant Proc, 1992, 24, 952. Jain K, Asina S, Yang H, Blount E D, Smith B H, Diehl C H and Rubin A L, ‘Glucose control and long-term survival in biobreeding/Worcester rats after intraperitoneal implantation of hydrophilic macrobeads containing porcine islets without immunosuppression’, Transplantation, 1999, 68(11), 1693–1700. Juang J H, Bonner-Weir S, Vacanti J P and Weir G C, ‘Outcome of subcutaneous islet transplantation improved by a polymer device’, Transplant Proc, 1995, 27, 3215– 3216. Juang J H, Bonner-Weir S, Wu Y J and Weir G C, ‘Beneficial influence of glycemic control upon the growth and function of transplanted islets’, Diabetes, 1994, 43(11), 1334–1339. Kessler L, Legeay G, Jesser C, Damgé C and Pinget M, ‘Influence of corona surface treatment on the properties of an artificial membrane used for Langerhans islets encapsulation: Permeability and biocompatibility studies’, Biomaterials, 1995, 16, 185–191. Kessler L, Legeay G, West R, Belcourt A and Pinget M, ‘Physicochemical and biological

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studies of corona-treated artificial membranes used for pancreatic islets encapsulation: Mechanism of diffusion and interface modification’, J Biomed Mater Res, 1997, 34, 235–245. Kessler L, Pinget M, Aprahamian M, Dejardin P and Damge C, ‘In vitro and in vivo studies of the properties of an artificial membrane for pancreatic islet encapsulation’, Horm Metab Res, 1991, 23, 312–317. Klomp G F, Ronel S H, Hashiguchi H, D’Andrea M and Dobelle W H, ‘Hydrogels for encapsulation of pancreatic islet cells’, Trans Am Soc Artif Intern Organs, 1979, 25, 74–76. Knazek R A, Gullino P M, Kohler P O and Dedrick R L, ‘Cell culture on artificial capillaries: an approach to tissue growth in vitro’, Science, 1972, 178, 65–66. Koo J and Chang T S M, ‘Secretion of erythropoietin from microencapsulated rat kidney cells’, Int J Artif Organs, 1993, 16, 557–560. Krause T J, Robertson F M, Liesch J B, Wasserman A J and Grecos R S, ‘Differential production of IL-1 on the surface of biomaterials’, Arch Surg, 1990, 125, 1158–1160. Kurian P, Kasibhatla B, Daum J, Burns C A, Moosa M, Rosenthal K S and Kennedy J P, ‘Synthesis, permeability and biocompatibility of tricomponent membranes containing polyethylene glycol, polydimethylsiloxane and polypentamethylcyclopentasiloxane domains’, Biomaterials, 2003, 24(20), 3493–3503. Lacík I, ‘Polymer chemistry in diabetes treatment by encapsulated islets of langerhans: review to 2006’, Aust J Chem, 2006, 59, 508–524. Lacy P E, Hegre O D, Gerasimidi Vazeou A, Gentile F T and Dionne K E, ‘Maintenance of normoglycemia in diabetic mice by subcutaneous xenografts of encapsulated islets’, Science, 1991, 254, 1782–1784. Lanza R P, Beyer A M and Chick W L, ‘Xenogeneic humoral responses to islets transplanted in biohybrid diffusion chambers’, Transplantation, 1994, 57, 1371–1375. Lanza R P, Borland K M, Lodge P, Carretta M, Sullivan S J, Muller T E, Solomon B A, Maki T, Monaco A P and Chick W L, ‘Pancreatic islet transplantation using membrane diffusion chambers’, Transplant Proc, 1992, 24, 2935–2936. Lanza R P and Chick W L, ‘Immunoisolation: at a turning point’, Immunology Today, 1997a, 18, 135–139. Lanza R P and Chick W L, ‘Transplantation of encapsulated cells and tissues’, Surgery, 1997b, 121, 1–9. Lanza R P and Chick W L, ‘Encapsulated Cell Therapy’, Scientific American Science and Medicine, 1995, 4, 16–25. Lanza R P, Ecker D M, Kühtreiber W M, Marsh J P and Chick W L, ‘Simple and inexpensive method for transplanting xenogeneic cells and tissues into rats using alginate gel spheres’, Transplant Proc, 1995, 27, 3322. Lanza R P, Hayes J L and Chick W L, ‘Encapsulated cell technology’, BioTechnology, 1996, 14, 1107–1111. Lanza R P, Kuhtreiber W M, Ecker D M, Marsh J P, Staruk J E and Chick W L, ‘A simple method for xenotransplanting cells and tissues into rats using uncoated alginate microreactors’, Transplant Proc, 1996, 28, 835. Lanza R P, Lodge P, Borland K M, Carretta M, Sullivan S J, Beyer A M, Muller T E, Solomon B A, Maki T, Monaco A P and Chick W L, ‘Transplantation of islet allografts using a diffusion-based biohybrid artificial pancreas: long-term studies in diabetic, pancreatectomized dogs’, Transplant Proc, 1993, 25, 978–980. Lanza R P, Sullivan S J and Chick W L, ‘Perspectives in diabetes. Islet transplantation with immunoisolation’, Diabetes, 1992, 41, 1503–1510.

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Lhommeau C, Toillon S, Pith T, Kessler L, Jesser C and Pinget M, ‘Polyamide 4,6 membranes for the encapsulation of Langerhans islets: preparation, physico-chemical properties and biocompatibility studies’, J Mater Sci Mater Med, 1997, 8(3), 163– 174. Lim F and Sun A M, ‘Microencapsulated islets as bioartificial endocrine pancreas’, Science, 1980, 210, 908–910. Liu H W, Ofosu F A and Chang P L, ‘Expression of human factor IX by microencapsulated recombinant fibroblasts’, Hum Gene Ther, 1993, 4, 291–301. Loudovaris T, Jacobs S, Young S, Maryanov D, Brauker J and Johnson R C, ‘Correction of diabetic nod mice with insulinomas implanted within Baxter immunoisolation devices’, J Mol Med, 1999, 77(1), 219–222. Maki T, Lodge J P, Carretta M, Ohzato H, Borland K M, Sullivan S J, Staruk J, Muller T E, Solomon B A, Chick W L and Monaco A P, ‘Treatment of severe diabetes mellitus for more than one year using a vascularized hybrid artificial pancreas’, Transplantation, 1993, 55, 713–717. Maki T, Mullon C J, Solomon B A and Monaco A P, ‘Novel delivery of pancreatic islet cells to treat insulin- dependent diabetes mellitus’, Clin Pharmacokinet, 1995, 28, 471–482. Maki T, O’Neil J J, Porter J, Mullon C J P, Solomon B A and Monaco A P, ‘Long-term function of porcine islets in xenogeneic hosts’, Transplant Proc, 1996a, 28, 807. Maki T, Otsu I, O’Neil J J, Dunleavy K, Mullon C J P, Solomon B A and Monaco A P, ‘Treatment of diabetes by xenogeneic islets without immunosuppression – Use of a vascularized bioartificial pancreas’, Diabetes, 1996b, 45, 342–347. Piro S, Lupi R, Dotta F, Patane G, Rabuazzo M A, Marselli L, Santangelo C, Realacci M, Del-Guerra S, Purrello F and Marchetti P, ‘Bovine islets are less susceptible than human islets to damage by human cytokines’, Transplantation, 2001, 71(1), 21–26. Ponce S, Orive G, Hernandez R, Gascon A R, Pedraz J L, De Haan B J, Faas M M, Mathieu H J and De Vos P, ‘Chemistry and the biological response against immunoisolating alginate-polycation capsules of different composition’, Biomaterials, 2006, 27(28), 4831–4839. Qi M, Gu Y, Sakata N, Kim D, Shirouzu Y, Yamamoto C, Hiura A, Sumi S and Inoue K, ‘PVA hydrogel sheet macroencapsulation for the bioartificial pancreas’, Biomaterials, 2004, 25(27), 5885–5892. Remes A and Williams D F, ‘Immune response in biocompatibility’, Biomaterials, 1992, 13(11), 731–743. Robitaille R, Dusseault J, Henley N, Desbiens K, Labrecque N and Halle J P, ‘Inflammatory response to peritoneal implantation of alginate-poly-L-lysine microcapsules’, Biomaterials, 2005, 26(19), 4119–4127. Sandler S, Eizirik D L, Sternesjo J and Welsh N, ‘Role of cytokines in regulation of pancreatic B-cell function’, Biochem Soc Trans, 1994, 22, 26–30. Scharp D W, Mason N S and Sparks R E, ‘Islet immuno-isolation: the use of hybrid artificial organs to prevent islet tissue rejection’, World J Surg, 1984, 8, 221–229. Scharp D W, Swanson C J, Olack B J, Latta P P, Hegre O D, Doherty E J, Gentile F T, Flavin K S, Ansara M F and Lacy P E, ‘Protection of encapsulated human islets implanted without immunosuppression in patients with type I or type II diabetes and in nondiabetic control subjects’, Diabetes, 1994, 43, 1167–1170. Sefton M V, ‘The good, the bad and the obvious: Clemson Award for Basic ResearchKeynote Lecture’, Biomaterials, 1993, 14(15), 1127–1134. Shoichet M S, Winn S R, Athavale S, Harris J M and Gentile F T, ‘Poly(ethylene oxide)-

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grafted thermoplastic membranes for use as cellular hybrid bio-artificial organs in the central nervous system’, Biotech Bioeng, 1994, 43, 563–572. Siebers U, Zekorn T, Bretzel R G, Planck H, Renardy M, Zschocke P and Federlin K, ‘Histocompatibility of semipermeable membranes for implantable diffusion devices (bioartificial pancreas)’, Transplant Proc, 1990, 22, 834–835. Soon Shiong P, Heintz R E, Merideth N, Yao Q X, Yao Z, Zheng T, Murphy M, Moloney M K, Schmehl M and Harris M et al., ‘Insulin independence in a type 1 diabetic patient after encapsulated islet transplantation’, Lancet, 1994, 343, 950–951. Sun A M, Parisius W, Healy G M, Vacek I and Macmorine H G, ‘The use, in diabetic rats and monkeys, of artificial capillary units containing cultured islets of Langerhans (artificial endocrine pancreas)’, Diabetes, 1977, 26, 1136–1139. Suzuki K, Bonner-Weir S, Hollister-Lock J, Colton C K and Weir G C, ‘Number and volume of islets transplanted in immunobarrier devices’, Cell Transplant, 1998a, 7(1), 47–52. Suzuki K, Bonner-Weir S, Trivedi N, Yoon K H, Hollister-Lock J, Colton C K and Weir G C, ‘Function and survival of macroencapsulated syngeneic islets transplanted into streptozocin-diabetic mice’, Transplantation, 1998b, 66(1), 21–28. Tam S K, Dusseault J, Polizu S, Menard M, Halle J P and Yahia L, ‘Physicochemical model of alginate-poly-L-lysine microcapsules defined at the micrometric/nanometric scale using ATR-FTIR, XPS, and ToF-SIMS’, Biomaterials, 2005, 26(34), 6950– 6961. Tatarkiewicz K, Hollister-Lock J, Quickel R R, Colton C K, Bonner-Weir S and Weir G C, ‘Reversal of hyperglycemia in mice after subcutaneous transplantation of macroencapsulated islets’, Transplantation, 1999, 67(5), 665–671. Thu B, Bruheim P, Espevik T, Smidrod O, Soon-Shiong P and Skjak-Braek G, ‘Alginate polycation microcapsules. I. Interaction between alginate and polycation’, Biomaterials, 1996, 17(3), 1031–1040. Trivedi N, Steil G M, Colton C K, Bonner-Weir S and Weir G C, ‘Improved vascularization of planar membrane diffusion devices following continuous infusion of vascular endothelial growth factor’, Cell Transplant, 2000b, 9(1), 115–124. Trivedi N, Steil G M, Colton C K, Bonner-Weir S and Weir G C, ‘Improved vascularization of planar membrane diffusion devices following continuous infusion of vascular endothelial growth factor’, Cell Transplant, 2000a, 9(1), 115–124. Uludag H and Sefton M V, ‘Metabolic activity and proliferation of CHO cells in hydroxyethyl methacrylate-methyl methacrylate (HEMA-MMA) microcapsules’, Cell Transplant, 1993, 2(2), 175–182. Van Hoogmoed C G, Busscher H J and De Vos P, ‘Fourier transform infrared spectroscopy studies of alginate-PLL capsules with varying compositions’, J Biomed Mater Res, 2003, 67A(1), 172–178. Woodward S C, ‘How fibroblasts and giant cells encapsulate implants: considerations in design of glucose sensors’, Diabetes Care, 1982, 5, 278–281. Yoon K H, Quickel R R, Tatarkiewicz K, Ulrich T R, Hollister-Lock J, Trivedi N, BonnerWeir S and Weir G C, ‘Differentiation and expansion of beta cell mass in porcine neonatal pancreatic cell clusters transplanted into nude mice’, Cell Transplant, 1999, 8(6), 673–689.

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22 Hydrogels for spinal cord injury regeneration A . J . S A L G A D O and N . S O U S A, Life and Health Sciences Research Institute (ICVS), University of Minho, Portugal, and N. A. S I L V A, N . M . N E V E S and R . L . R E I S, 3B’s Research Group, University of Minho, Portugal

22.1

Introduction

Spinal cord injury (SCI) affects millions of individuals worldwide. The US and EU alone have reported around 30 000 annual cases of unrecoverable SCI. Due to recent improvements in the field, the life expectancy of affected patients has been increasing year by year, a fact that could lead in 10 years to a total of 600 000 affected people in both the US and EU (Berkowitz et al., 1992). However, these patients still suffer from common problems related with SCI such as frequent infections in bladder, kidneys, bowel problems and eschars. Furthermore it is also frequent that SCI afflicted individuals develop emotional/mood disorders such as depression. In addition there are also significant reductions in rates of occupation and employment after injury, a fact that also contributes to the last point that was raised. In parallel to these social problems, SCI also represents a tremendous burden to community and health care systems worldwide. The costs associated with SCI are mainly related to the initial and subsequent hospitalizations, rehabilitation and supportive equipment, home modifications, personal assistance, institutional care and loss of income (Samadikuchksaraei, 2007). For instance the average lifetime cost that is directly attributed to SCI is estimated to be US$620 000– $2 800 000 for patients aged 25 and US$450 000–$1 600 000 for each patient aged 50 at the time of injury (Samadikuchksaraei, 2007; Sekhon and Fehlings, 2001). In recent years an increasing number of strategies aimed at adequately tackling this problem have been reported. Among these, biomaterial and tissue engineering strategies have been at the forefront of a new wave of hybrid strategies that try to deal with the problem of SCI regeneration, in an integrative manner. The present chapter will be focused on the latter, namely those strategies based on the use of hydrogels. It will start by giving an overview of the structural components of the central nervous system (CNS) and, within it, the spinal cord followed by a brief description on the therapies currently used in the field, as well as other innovative strategies that are currently being used for SCI repair, such as cell based therapies. The objective 570 © 2008, Woodhead Publishing Limited

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of the latter is to give some background to the readers on basic topics of CNS structure and cell biology, so they can fully understand the topics that will be covered in the hydrogel section. The latter will then be mainly focused on the most used hydrogels within the field, from synthetic or natural origins, either used alone, loaded with growth factors or associated with cell populations with the capability of inducing SCI regeneration. The chapter will then be finalized by taking a forward look at, what are in the authors’ opinion, the future trends within the SCI field.

22.2

Brief insights on central nervous system biology

Since the beginning of time the nervous system has intrigued human civilization. For instance Plato used to call the brain our ‘Divine Area’ (Plato, 1993). Presently, and in spite of deep study, it still encases many mysteries, and therefore its pathologies still represent a major challenge for biomedical science. In a quite simplistic form the nervous system gathers, processes and transmits stimuli/impulses, either physical or chemical, in order to allow the human body to adapt itself to the environment that surrounds it, allowing in this sense the maintenance of the homeostasis within the human body (Fox, 2003).

22.2.1 The central nervous system – basic concepts The nervous system is structurally divided into two sections (Fox, 2003; Junqueira, 2004): (1) central nervous system (CNS); and (2) peripheric nervous system (PNS). The first is constituted by the brain and spinal cord, and is considered to be the ‘command post’ of the nervous system. It is responsible for the interpretation and processing of information that is handed over by the PNS, which is constituted by an array of nerves that are spread throughout the body and connected to the spinal cord. Its main function is to gather information from the outer/inner environment. In this sense the spinal cord can be looked as an information ‘highway’ transmitting information from and into the brain, therefore playing a leading role in the normal functioning of the CNS. From the cell biology point of view the CNS is composed of two cell types, neurons and glial cells. Neurons (Figure 22.1a) are considered to be the structural and functional unit of the CNS. Their main characteristic is based on the ability to respond and conduct electrochemical impulses, their actions being mainly mediated through the release of chemical regulators known as neurotransmitters (Junqueira, 2004; Marieb, 1992). Although neurons can considerably change in shape and size, they are normally divided into three areas: the cell body or soma, dendrites and the axon (Junqueira, 2004). The cell body, as its

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22.1 Fluorescence microscopy micrographs of: (a) Hippocampal neurons; (b) Astrocytes; (c) Oligodendrocytes and (d) Microglial cells.

name implies is the area where all the nucleous and the remaining organelles are located, and therefore is considered as the ‘nutritional centre’ of the neuron, as it is the place where macromolecules found to be essential to neuronal survival are produced (Marieb, 2004). The dendrites can be looked upon as extensions of the cell body that provide an extended surface area for the reception of the nervous impulses, transmitted by neighbouring neurons, to the cell bodies (Marieb, 1992). Finally, the axon, can also be looked upon as an extension protruding from the cell body, although much more differentiated than the dendrites. It is through this structure that the nervous impulses are conducted from the cell body to the axon terminals, where neurotransmitters will be released to adjacent neurons to an area denominated as synaptic cleft (Junqueira, 2004; Marieb, 1992). The axon is normally surrounded by a lipoproteic sheath of myelin. The latter acts mainly as an electric insulator, allowing in this sense a faster transmition of the nervous impulse. The areas of the brain that are rich in myelinated fibers are known as white matter while the areas containing the cell bodies are known as grey matter (Junqueira, 2004, Marieb, 1992). The glial cells are the second cell type of the central nervous system. They are more abundant than the neurons and, unlike them, they do have

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mitototic properties (Junqueira, 2004). Glial cells are mainly involved in supporting functions such as the maintenance of the ionic balance of neurons, the modulation of the synaptic transmission and the re-establishment of the neurological function upon neuronal damage (although this is not always possible, as in the case of SCI). Three cell populations can be found among glial cells (Junqueira, 2004): astrocytes, oligodendrocytes and microglial cells. Astrocytes Astrocytes (Figure 22.1b) are the most abundant glial cell type. They are commonly located in the vicinities of blood vessels and neurons, and are, among others, responsible for mediating the synaptic transmission by removing, from the synaptic cleft, neurotransmitters that will be later ‘recycled’ and made available again to neurons. Moreover, they are also deeply involved on nutrient exchange with neurons and simultaneously they make part of the blood brain barrier (BBB), which can be looked on as a barrier against foreign agents that intend to enter the CNS (Marieb, 1992). Oligodendrocytes Oligodendrocytes (Figure 22.1c) are the cells involved in the production of the myelin sheath that encases the neurons. Therefore they are of the utmost importance within the CNS, as the loss of myelinization leads to extensive damage within the latter, as happens during SCI (Fox, 2003). Microglial cells Microglial cells (Figure 22.1d) play an important role within the CNS, as they are considered to be the immune mediators of the latter, possessing a macrophage like behaviour, phagocyting death tissue and cells, as well as foreign agents. They possess two states, which mainly reflect their state of activation. Ramified microglial cells are considered to be in the resting state. They are characterized by possessing long cytoplasmastic processes, a small cell body and few lysosomes. On the other hand, activated microglia are characterized by ameboid morphology, with a large cell body and a large number of lysosomes, which indicates a high phagocytic capacity (Purves, 1997).

22.2.2 The spinal cord – anatomy and motor circuits In a rather simplistic manner, the spinal cord (Figure 22.2) can be looked on as a strand of tissue that begins at the base of the cerebellum, going thereafter through the magnum (an orifice on the base of the skull) until it reaches the

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22.2 The spinal cord: (a) Spinal reflex: (1) spinal cord; (2) grey matter; (3) white matter; (4) interneuron; (5) dorsal root ganglion; (6) dorsal root; (7) sensory neuron; (8) effecter muscle; (9) somatic motor neuron; (10) spinal nerve; (11) ventral root; (b) Vertebral body: (1) spinal nerve; (2) pia mater; (3) subarachnoid space; (4) arachnoid; (5) dura mater; (6) epidural space; (7) spinal cord; (8) vertebral body; (c) Organization of the spinal cord: (1) anterior (ventral) root; (2) posterior (dorsal root); (3) spinal nerve; (4) posterior (dorsal) root ganglion; (5) white matter; (6) grey matter; (7) central canal. (Adapted from Fox 2003; Mader 2004; Van de Graaf 2001.)

first lumbar vertebrae. It is divided into four regions (Netter, 2003): (1) cervical, which possesses eight nerves and coordinates the upper portion of the trunk; (2) thoracic, which controls the lower portion of the trunk and encases 12 nerves; (3) lumbar, which possesses five nerves and coordinates the movement of the lower limbs; and finally (4) sacral, composed of six nerves, controlling the genital areas, as well as the bowel and bladder functions. The spinal cord has two main functions: (1) to conduct the nervous impulses, through the ascending pathways/nerves, which mainly deal with sensorial stimuli, and the descending nerves, dealing with the motor function, establishing in this sense a network between the rest of the body and the brain; and (2) integrate the reflexes, being able to induce simple movements as involuntary reflexes, such as sneezing and vomiting. Like the brain, the spinal cord is protected by a bony structure, the vertebral column, layers of connective tissue that form membranes known as meninges, namely the dura mater, the arachnoide and the pia mater and by cerebrospinal

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fluid. Within it, the grey matter is placed in the centre, being surrounded by the white matter. The grey matter has an ‘H’ like shape and within its structure one can find motor neurons and inter-neurons, while the surrounding white matter encases the ascending and descending nerves (Van der Graaf, 2001).

22.2.3 Spinal cord injury – from acute to chronic stage SCI is usually caused by blunt impacts, compression episodes or penetrating traumas, such as gunshot incidents or stab wounds. Blunt impacts usually lead to concussion, contusion, laceration, transaction or intraparenchymal haemorrhage, while spinal cord compression usually results from hyperflexion, hyperextension, axial loading and severe rotation (Samadikuchaksaraei, 2007). The damage that is caused at the time of the trauma to the spinal cord is called the ‘primary injury’, which then develops into a cascade of events known as secondary injury. Several mechanisms are involved in secondary injury, vascular changes being the most important events. The microvascular alterations include loss of autoregulation, thrombosis and haemorrhage, and, when associated with haedema lead to hypoperfusion, ischaemia and necrosis (Abraham et al., 2000; Samadikuchaksaraei, 2007). Other major mechanisms include (Abraham et al., 2001, Barami and Diaz, 2000; Conti A et al., 2003; Houle and Tessler, 2003; Knoblach et al., 2005; Ray et al., 2003; Schwab 2002, Schwab, 2004; Takagi et al., 2003; Yang and Pias, 2003): (1) free radical formation and lipid peroxidation; (2) accumulation of excitatory neurotransmitters which causes neural damage due to excitotoxicity; (3) loss of intracellular balance of sodium, potassium, calcium and magnesium, leading to an increase in the levels of intracellular calcium; (4) higher concentration of opioids, namely dynorphins, at the site of injury; (5) depletion of energy metabolites leading to anaerobic metabolism at the site of injury and increasing LDH activity; (6) onset of an inflammatory response with the consequent recruitment and activation of inflammatory cells associated with secretion of cytokines, contributing in this sense to further tissue damage; and (7) activation of calpains, caspases and apoptosis. Furthermore SCI leads to other events that are predominantly deleterious to local cell populations and nerve tracts. For instance, penetrating injuries commonly lead to extensive oligodendrocyte cell death, causing dyemilination, and consequent impairment of nervous impulse conduction (Filbin et al., 2003). Contusion injuries are also characterized by the formation of a cystic cavity surrounded by an astrocytic scar, which acts as a physical barrier that does not allow the axons to grow across the cavity (David and Hacroix, 2003). Finally, in both crushed and transacted nerve fibres it is also common to find inhibitory proteins such as Nogo, MAG and chondroitin sulphate proteoglycans (CSPGs) that further block regenerative sprouting (Samadikuchaksaraei, 2007).

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22.3

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Current approaches for SCI repair

Upon injury, the spinal cord is stabilized and its original alignment restored, together, in most cases, with decompression of the cord. Following this the most common treatments are based on the use of pharmacological agents, which are more oriented towards neuroprotection, or alternatively, the use of experimental techniques such as cell based therapies. The following subsections contain brief descriptions of both topics.

22.3.1 Pharmacological approaches The use of pharmacological agents in the context of SCI is under the pretence that even small gains in neuroprotection might affect functional relevant neurological recovery (Kwon, et al., 2004). Corticosteroids Corticosteroids are among the most classical approaches and treatments in the context of SCI, the mostly used one being methylprednisolone. The precise mechanisms by which corticosteroids affect neuroprotection are not completely understood but are believed to include the inhibition of lipid peroxidation and inflammatory reactions, modulation of the inflammatory/ immune cells, improved vascular perfusion and prevention of calcium influx and accumulation (Kwon et al. 2004, Young 2000). Opioid antagonists The use of opioid antagonists has also been a matter of intense research. For instance, the non-specific opioid receptor antagonist naloxone was intensively investigated in the early 1980s after it was found to reverse spinal shock and improve spinal cord blood flow, with associated functional and electrophysiologic improvements in animal models of SCI (Holaday and Faden 1980; Faden et al., 1981a, b; Flamm et al. 1982; Young et al., 1981). Further studies revealed that naxolone did in fact promote motor and sensory recovery in incompletely injured patients (Bracken and Holford 1993). Glutamate receptors and ion channel antagonist have also been suggested, namely those dealing with NMDA and non-NMDA (AMPA/Kainate) receptor activation, as their role has been recognized in excitotoxic damage after SCI (Kwon et al., 2004). For instance NMDA receptor antagonists MK801 and gacyclidine (GK11), have demonstrated significant neuroprotective effects after experiments carried out in SCI animal models (Gaviria et al., 2000a, b). However the development of clinical therapies using the above referred strategies have been somewhat refrained due to the widespread distribution of glutamate (and its receptors) in neurotransmission throughout the human CNS (Kwon et al., 2004). © 2008, Woodhead Publishing Limited

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Others Other strategies used in this field have been the application of nonsteroidal anti-inflammatory agents, such as ibuprofen and meclofenamate. This application is based on the control of inflammatory prostaglandins in SCI chronic injury, namely through the use of cyclooxygenase inhibitors, as is the case of the above referred agents (Hall and Wolf, 1986). Besides these, other pharmacological agents are currently being tested in animal models. Among those, minocycline, a tetracycline antibiotic, has been shown to inhibit excitotoxicity. It is currently under investigation in animal models of contusive spinal cord injury, with preliminary results suggesting a promising reduction in tissue damage and apoptotic death at the injury site, as well as improved locomotor function (Arnold et al., 2001; Tikka et al., 2001).

22.3.2 Cell based therapies One of the experimental techniques that have been proposed throughout the years for SCI regeneration is the injection of different cell populations within the affected areas. Due to their characteristics, Schwann cells were one of the first cell populations to be proposed for SCI regeneration. These are the myelinating cells from the peripheral nervous system, and this is the purpose of using them in CNS related injuries. Furthermore they are commonly one of the first cell types to respond to injury within the spinal cord (Guest et al., 2005; Jasmin et al., 2000). It has been previously shown that these cells are able to promote axonal regeneration through the secretion of adhesion molecules such as L1 and N-CAM, ECM molecules such as collagen and laminin, and a number of trophic factors such as FGF-2, NGF, BDNF and NT-3 (Chernousov and Carey, 2000; Chernousov et al., 2000; Grothe et al., 2001; Guest et al., 2005; Jasmin et al., 2000; Kohama et al., 2001; Pinzon et al., 2001). Another cell population with myelinating capability that has been used in SCI related research are olfactory ensheathing cells (OECs) (Ramón-Cueto and Nieto-Sampedro 1994). These cells are considered to be glial cells, possessing properties of both Schwann cells and astrocytes, and can be isolated from the olfactory bulb or the nasal mucosa (lamina propria) (Samadikuchaksaraei, 2007). OECs isolated from these sources have been shown to induce regeneration and functional recovery after transplantion to animal models of transection, hemisection and contused spinal cords (Au et al., 2007; Sasaki et al. 2006; Ramón Cueto et al., 2000). A more recent trend in the field has been the use of adult stem cells, namely derived from the CNS, neural stem cells (NSCs) and the bone marrow (haematopoietic stem cells – HSCs; mesenchymal stem cells – MSCs). NSCs are mainly found within the subventricular (SVZ) zone of the brain and have

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previously been shown to differentiate towards the neuronal and glial lineages (Reynolds et al., 1992). It is known that when transplanted into neurogenic sites these cells can differentiate into the above referred lineages (Merkle et al., 2007). However, when injected into non-neurogenic sites, such as the spinal cord, they mostly differentiate into glial cells, namely astrocytes (Ogawa et al., 2002; Iwanami et al., 2005). Still, it has been shown in a number of studies that these cells lead to axonal regeneration and functional improvements when injected into SCI sites (Su et al. 2007; Zhang et al., 2007). HSCs and MSCs are not from neural origin, but recent reports have shown their possible application to the SCI field. In both cases their transplantation to injury sites led to an overall improvement in the condition of the animals (different models of SCI). Although it has been previously reported that both cell populations had the capacity to differentiate into neurons and glial cells, even if the yields of such phenomena were very low, it is most likely that the observed effects are related to the known capability of these cells to release trophic factors and neurogulatory molecules to the extracellular environment (Deng et al., 2006; Koshizuka et al., 2004).

22.4

Hydrogel-based systems in SCI regenerative medicine

Although the above mentioned strategies, namely those based on cell therapies, have shown some promising results, one must not forget that within SCI sites there are a plethora of molecules and physical barriers that most of the time inhibit axonal growth. It is believed that this happens due to the lack of a support matrix at the lesion site to direct the migration and organization of local wound-healing cells, as well as to promote directed axonal outgrowth across the repaired lesion (Woerly et al., 2001a). The need for a support matrix at the site of injury is further substantiated by the fact that even when ischaemic necrosis, cavitation and scar formation are limited or absent, no cellular or axonal ingrowth occur in the lesion site (Guth et al., 1981). In this sense there is a need to direct localized cellular regeneration events by creating a suitable environment for migrating tissue repair cells and regenerating axons (Woerly et al., 2001b). This can be achieved by introducing a threedimensional substrate into the site of injury, which will then provide a structural continuity across the lesion (Woerly et al., 1999; Woerly et al., 2001b). This concept arises from studies on the role of extracellular matrices (ECM) which provide a skeletal tissue framework in the construction of tissue structure during development, remodelling and regeneration (Adams and Watt, 1993). ECM form macromolecular networks of glycoproteins and heteropolysaccharides which, in living tissues, form highly hydrated gel-like structures due to interactions with extracellular fluids. The latter will then act as a hydrated environment for cell proliferation and supracellular

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organization, and for diffusion of neuroactive factors essential for cell survival and differentiation. One possible form of obtaining these so called spinal cord ‘bridges’ could be through the transplantation of tissue grafts from, for instance, the peripheric nervous system. This has already been done, but with various rates of success (Kuo et al., 2007; Rasouli et al., 2006). However the usual problems associated with tissue grafts are presented, such as the limited amount of tissue available (Salgado et al., 2004). It is in this context that biomaterial based approaches, and within it hydrogels, have been put forward as a possible alternative for the development of strategies in SCI regenerative medicine. Hydrogels consist of a crosslinked network of hydrophilic polymers that swell in water, culture medium or biological fluids providing in this sense an adequate environment for cell attachment and growth (Woerly et al., 2001b). Furthermore, and because of their ability to retain substantial amounts of water with respect to network density, hydrogels allow the transport of small molecules, such as growth factors. In addition, viscoelastic behaviour, low interfacial tension with biologic fluids and structure stability make porous hydrogels suitable for implantation in SCI affected sites (Woerly et al., 2001b). Moreover due to their characteristics, hydrogels are also amenable for cell encapsulation, which will certainly be of use, and will definitely help in the healing and regenerative processes of the SCI affected site. As for other classes of biomaterials, hydrogels can be from synthetic or natural origin. Both of them present advantages and disadvantages, but have shown promising results (Woerly et al., 2001b). In the following sections we will present and discuss those that are believed to be more relevant to the field.

22.4.1 Synthetic based hydrogels Synthetic materials have many advantages for use as regenerative bridges in SCI regeneration. These polymers can be tailored to produce a wide range of mechanical properties and degradation rates. They also have known compositions and can be designed to minimize the immune response. Finally, synthetic polymers can be reacted together to combine the properties that are unique to each (Willerth and Sakiyama-Elbert, 2007). Poly [N-2-(hydroxypropyl) methacrylamide] (PHPMA) Poly [N-2-(hydroxypropyl) methacrylamide] (PHPMA) based hydrogels were first described by Woerly and colleagues (Woerly et al., 1998; Woerly et al. 1999). They are characterized by possessing a multimode pore distribution, comprising micropores (< 2 nm), mesopores (2–50 nm) and macropores (50–300 nm). The micropores allow the transport of small molecules, mesopores

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the transport of larger molecules, while macropores allow both free transport of large molecules, as well as the migration of cells and blood vessels (Woerly et al., 1999). Furthermore the diffusion properties of this hydrogel allow it to maintain a physiological environment in the implantation site, facilitating diffusion of neurotrophic molecules that are released from the reactive cells of host CNS tissues (Woerly et al., 1999). According to the authors, other important features are its viscoelastic characteristics, its low interfacial tension for biological fluids and its structural stability (Woerly 2000; Woerly 2001a; Woerly, 2001b). When placed in SCI areas these PHPMA hydrogels seem to improve regeneration as well as motor function (Woerly, 2000; Woerly, 2001a; Woerly, 2001b). Poly(2-hydroxyethyl methacrylate-co-methyl methacrylate) (pHEMA-MMA) Poly(2-hydroxyethyl methacrylate-co-methyl methacrylate) (PHEMA-MMA) is another example of a non-biodegradable synthetic hydrogel that has been under study in SCI regenerative medicine (Tsai et al., 2004; Tsai et al., 2006). Due to their non-degradable nature they are able to remain stable upon implantation (Willerth and Sakiyama-Elbert 2007). Furthermore, due to their versatility they can be moulded into a variety of shapes, which might be of help for their application in clinical practice. In a study published by Tsai et al., (2004) the biocompatibility and regenerative capacity of synthetic pHEMA-MMA hydrogels was assessed in a rat transaction model. Gross and microscopic examinations of the spinal cord showed continuity of tissue within the synthetic guidance channels between the cord stumps at four and eight weeks. In another study published by the same authors (Tsai et al., 2006) a modification was introduced to the pHEMA-MMA hydrogel. In order to improve its overall characteristics, as well as the number and type of regenerated axons within the spinal cord, different matrices and growth factors were incorporated within the lumen of the pHEMA-MMA guidance conduit. After complete spinal cord transection at T8, pHEMA-MMA channels were implanted into adult Sprague Dawley rats. The conduits had been previously filled with one of the following matrices: collagen, fibrin, Matrigel, methylcellulose, or smaller pHEMA-MMA tubes placed within a larger pHEMA-MMA channel (called tubes within channels, TWC). Collagen and fibrin matrices were additionally supplemented with neurotrophic factors: fibroblast growth factor-1 (FGF-1) and neurotrophin-3 (NT-3). Results showed that fibrin, matrigel, methylcellulose, collagen with FGF-1, collagen with NT-3, fibrin with FGF-1, and fibrin with NT-3 increased the total axon density within the channel compared to unfilled channel controls. This study showed that these pHEMA-MMA hydrogel based tubes could be quite versatile, through the combination with other hydrogels in which future encapsulation

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of relevant cell populations could be envisaged. However further studies need to be done, namely regarding the long term structural integrity of these guidance conduits. For instance, in a report by Belkas et al. (2005) it was shown that after 16 weeks of implantation the degree of integrity of the conduits was drastically reduced when compared to the eight weeks’ time point. Poly (ethylene glycol) Polyethylene glycol (PEG) is a water-soluble surfactant polymer. It is characterized by having a resistance to protein adsorption, which might help to minimize immune response upon in vivo implantation (Willerth and Sakiyama-Elbert, 2007). Hydrophilic PEG hydrogels can be made through a variety of cross-linking reactions to create guidance structures with varying degradation rates (Burdick et al., 2006; Elbert and Hubbell, 2001, Mahoney and Auseth, 2007; Nguyen and West, 2002; Peppas et al., 2002). Brief application of an aqueous solution of this polymer to the site of injury in the spinal cord seals and repairs cell membrane breaches, reverses the permeabilization of the membrane produced by injury, inhibits production of free radicals (Luo et al., 2002; Shi and Borgens, 1999, Shi and Borgens, 2000) and decreases oxidative stress (Luo and Shi, 2004; Luo et al., 2004). It has been previsouly shown that PEG is able to re-establish the anatomical continuity and lead to functional recovery of severed rat guinea pig spinal cord (Piantino et al. 2006, Shi et al. 1999).

22.4.2 Natural-based hydrogels Natural materials possess many properties that make them attractive for neural tissue applications. Many of these materials contain sites for cell adhesion and cell infiltration. These materials also exhibit similar properties to the soft tissues they are replacing. Since these materials are obtained from natural sources, they must be purified to ensure that no foreign body response occurs after implantation. Homogeneity of product between lots can be an issue with natural materials (Willerth and Sakiyama-Elbert, 2007). Alginate Alginate is a naturally derived polysaccharide extracted from brown sea weed and composed of 1,4-linked β-D-mannuronate and 1,4-linked α Lglucuronate. It can be cross-linked by complexation of its carboxylic groups with many multivalent cations such as Cu2+, Ca2+ or Al3+, thus producing mechanically stable hydrogels (Prang et al., 2006). Calcium alginates represent biocompatible and non-immunogenic polymers, which have been used as

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scaffold material for tissue engineering and transplantation of cells (Prang et al., 2006). As the calcium ions gradually diffuse out from the gel, the latter slowly degrades and it is excreted in the urine (Novikova et al., 2006). The degradation rate mainly depends on the concentration of alginate, dose of irradiation, and cross-linked density (Suzuki et al., 2002). Furthermore both in vitro and in vivo models have previously demonstrated that alginates are biodegradable without causing allergic/inflammatory reactions (Orive et al., 2002; Prang et al. 2006; Suzuki et al., 1999; Suzuki et al., 2002). However it must be noticed that in order to achieve such properties the alginate based biomaterials must undergo extensive purification to prevent immune responses after implantation (Willerth and Sakiyama-Elbert, 2007). Regarding SCI regeneration, alginate based scaffolds have previously been shown to possess some interesting properties for these purposes. For instance, Suzuki and colleagues (Suzuki et al., 1992) demonstrated that freeze-dried alginate sponges were able to induce an overall improvement in a rat spinal cord transection model. Results revealed that regenerated axons were able to traverse the alginate-filled gap, re-enter the host spinal cord, and form neural connections with host neurons. Both ascending and descending axons were considered to form polysynaptic connections with the GN and lumbar motorneurons, respectively. Furthermore electrophysiological data indicated that both ascending and descending regenerating axons had the capacity to form functional synapses with neurons beyond the lesion, which is a clear signal of functional improvement. However, in the same study it was also observed that the axonal regeneration was somewhat limited, a fact that was later attributed by Prang et al. (2006) to the fact that axonal regrowth rarely occured in a longitudinally oriented fashion, thus preventing reconnection with the caudal spinal cord. In order to overcome this problem these authors put forward a novel concept within the alginate based hydrogels for SCI repair, which was the development of alginate-based anisotropic capillary hydrogels (ACH) (Prang et al., 2006). The latter were based on the concepts previously reported by Thiele (1967), which stated that ACH are formed when an aqueous solution of sodium alginate and a solution containing multivalent cations are superimposed in layers under the assumption that convection is avoided. When placed on a transaction model of SCI, at the C3 level, alginate-based ACH were able to integrate into the spinal cord parenchyma without major inflammatory responses, maintain their anisotropic structure and direct axon regeneration across the artificial scaffold. Furthermore, in supplementary experiments they have also been shown to promote adult neural progenitor cell (NPC) proliferation and differentiation, which might be of use in future approaches, namely those that are focused on the transplantation of constructs containing adequate cell populations for SCI regeneration (Prang et al., 2006). In fact this alginate-based ACH seems to represent some progress within the alginate field, as a previous report by

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Novikova et al., (2006) showed that when OECs, SCs and BMSCs were encapsulated within standard alginate hydrogels they somewhat changed their morphology into atypical cells with spherical shape and inhibited metabolic activity. Agarose Agarose is a thermosensitive natural polysaccharide polymer. Agarose hydrogels have been used as substrates for model 3-D cell culture studies, cell encapsulation, and tissue engineering (Bellamkonda et al., 1995; Kuroki et al., 2003; Miyoshi et al., 1996). Like other polysaccharide polymers, agarose can be conveniently modified through the activation of hydroxyl groups as has been demonstrated with several reagents including cyanogen bromide, glycidol, 4-nitrophenyl chlorocarbonate, 1,1-carbonyldiimidazole, among others (Guisan et al., 1997; Khon and Wilchek, 1981; Miron and Wilchek, 1987). In recent years Molly Shoichet’s lab has put forward a new class of agarose based hydrogels that have shown interesting properties for CNS and SCI regenerative medicine (Luo and Shi, 2004; Luo et al., 2004). The latter are based on the photoimmobilization of biomolecules (namely GRGDS peptides) on agarose based hydrogels. Further details on the processing methodology can be found in the reports by Luo and Shoichet (2004a). When the GRGDS peptide was immobilized on agarose, it was shown to be cell-adhesive and to promote neurite outgrowth from primary, embryonic chick dorsal root ganglion neurons. The immobilized GRGDS surface ligand concentration affected the cellular response: neurite length and density increased with GRGDS surface concentration at low adhesion ligand concentration and then plateaued at higher GRGDS concentration. Using the same methodology the same authors also developed an interesting concept, which was the formation of chemical GRDGS channels in previously defined areas of the agarose based hydrogels, allowing in this sense the selective and oriented growth of axons (Luo and Shoichet, 2004b). Matrigel Matrigel is an extracellular matrix extracted from the Engelbreth Holm Swarm (EHS) sarcoma and contains laminin, fibronectin, and proteoglycans (Freshney, 2000). It is commonly used as a substrate for in vitro cell growth, migration and differentiation assays. Several reports have shown its possible application in the SCI regenerative medicine field. Implantation of Matrigel alone does not increase regenerative activities in the spinal cord (Xiao et al., 2005). However when combined with various cell populations and growth factors it has been shown to induce SCI regeneration. For instance, Matrigel combined with vascular endothelial growth factor (VEGF) decreased retrograde

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degeneration of corticospinal tract axons and increased axonal regenerative activities in rats (Facchiano et al., 2002). Similar behaviour was also detected for glial cell line derived factor (GDNF) enriched matrigels constructs (Ianotti et al., 2003). Simultaneously the inclusion of Matrigel within hydrogel guidance channels increased the number of regenerating axons penetrating the construct (Tsai et al., 2006). Matrigel has also been used as a scaffold for in vivo delivery of different cell populations. Pinzon et al. (2001) reported that the use of Schwann cells with Matrigel increased the number myelinated axons, blood vessels, macrophages and fibroblasts in the ‘bridging site’. Electrophysiological studies further showed the functionality of regenerating axons. In a recent study it was also shown that Matrigel could support the growth of olfactory ensheathing cells (OECs) (Fouad et al., 2005). OECs were implanted to enable regenerating axons, and chondroitinase ABC was used to reduce the axonal regeneration inhibitory effect of chondroitin sulphate proteoglycan (CSPG) in the glial scar. This combined implantation therapy significantly increased the number of myelinated axons and serotonergic fibres in the bridge. Simulaneously a significant functional improvement was observed in the animal under study (Willerth and Sakiyama-Elbert, 2007). Collagen Collagen, one of the most common ECM proteins, has been extensively characterized as a potential scaffold for neural tissue engineering. Collagen can be isolated from mammals, including rats, bovines and humans. By changing the pH of collagen solutions, gel formation can be induced. They are natural materials but an immune response could arise if cross-species transplantation is used. These scaffolds contain sites for cell adhesion and can be covalently modified (Archibald et al., 1991; Ishii et al., 2006; Grothe et al., 2006). Their properties can be varied by using different concentrations of collagen. Jimenez Hamann et al. (2005) have recently reported a drug delivery system (DDS) consisting of a highly concentrated collagen solution which could be injected intrathecally at the site of injury providing localized delivery of growth factors. Using the injectable DDS, epidermal growth factor (EGF) and basic fibroblast growth factor (FGF-2) were co-delivered in the subarachnoid space of Sprague-Dawley rats. Significant differences in the distribution of EGF and FGF-2 in the spinal cord were evident. Localized delivery of the growth factors resulted in significantly less cavitation at the lesion epicentre and for at least 720 µm caudal to it compared to control animals without the DDS. There was also significantly more white matter sparing at the lesion epicentre in animals receiving the growth factors compared to control animals. Furthermore, at 14 days post-injection, delivery of the growth factors resulted in significantly greater ependymal cell proliferation in the central canal immediately rostral and caudal to the lesion edge compared

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to controls. Yoshii et al. (2004) have also shown that collagen based hydrogel filaments could improve the condition of SCI afflicted animals, namely in a rat contusion/transaction model. A progressive analysis throughout the time length of the experiments revealed that at early time points (four weeks), regenerated axons crossed the proximal and distal spinal cord-implant interfaces, while at late time points (12 weeks) the rats could walk and revealed hind-forelimb coordination. Simultaneously somatosensory-evoked potentials were also observed. Collagen gels have also been used in combination with cell based therapies, where different cell populations are encapsulated in collagen based hydrogels prior to implantation in the injury site. For instance Wang and colleagues (Wang et al., 2006) have shown that hydrogel like collagen based scaffolds could support the growth of OECs. Fibronectin Fibronectin (Fn) occurs in plasma as a dimeric folded globular protein, but when incorporated into the extracellular matrix it takes on an insoluble fibrous elongated form. The structure of Fn has been well characterized and it is known to comprise a series of repeats of three distinct types of independently folded modules. The precise mechanism by which Fn is assembled into the fibrillar matrix by cells has yet to be elucidated, but subjecting the molecule to tensile forces both through mechanical shearing at a moving plastic–liquid interface or by attachment to a lipid interface model has been shown to result in Fn self-assembly (Phillips et al., 2004). Following this principle a series of studies come out from Robert Brown and John Priestley’s labs where fibronectin based hydrogels were tested as scaffolds for SCI repair (King et al., 2003; King et al., 2006; Phillips et al., 2004) . Phillips et al. (2004) showed that based on the shear aggregation principles it was possible to develop fibronectin structures capable of supporting CNS and PNS derived cell growth, such as astrocytes and Schwann cells. Furthermore when these materials were inserted in a transection like spinal cord model, they provided a permissive environment for axonal growth. Using another strategy, King et al., (2003) developed fibronectin mats for the above referred purposes. The latter contained pores, ranging in size from 10 to 200 µm and oriented in a single direction, that provided conduits for regenerating axons as well as a means for manipulating the direction of axonal growth. These pores were formed between aligned fibres that provided a cell-adhesive substrate for ingrowing cells or axons. Fn mats also absorb growth factors, releasing them as the mat dissolves, thus acting as a reservoir for neurotrophic factors. These mats were later implanted in lesion cavities 1 mm in width and depth and 2 mm in length, created on one side of the spinal cord of adult rats. Fn mats containing neurotrophins or saline were placed into the lesion. The mats were well integrated into surrounding tissue

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and showed robust well-oriented growth of calcitonin gene-related peptide, substance P, GABAergic, cholinergic, glutamatergic and noradrenergic axons into them. Electron microscopy confirmed the presence of axons within implant sites, with most axons either ensheathed or myelinated by Schwann cells. Mats incubated in brain-derived neurotrophic factor and neurotrophin3 showed significantly more neurofilament-positive and glutamatergic fibres compared to saline- and nerve growth factor-incubated mats, while mats incubated with nerve growth factor showed more calcitonin gene-related peptide-positive axons. In a subsequent report (King et al., 2006) the authors identify some of the mechanisms that were involved in the regeneration of SCI through the use of fibronectin mats. According to the authors, implantation of Fn mats into the spinal cord is followed by a cascade of events, which leads to the infiltration of a number of cellular and extracellular elements that, in addition to replacing the mat as it dissolves, stimulate axonal growth. Specifically, the initial infiltration of macrophages is followed by the infiltration of Schwann cells, which in turn deposit laminin in the lesion site, with each of these elements potentially providing trophic support for ingrowing axons. Later, as laminin and macrophages begin to disappear and astrocytes begin to infiltrate, the implant continues to be occupied by myelinated axons. The replacement of the mat by these endogenous biological materials results in an implant site that is made up largely of myelinated axons. Fibrin Fibrin, which serves as the natural wound healing matrix that forms after injury and its precursor, fibrinogen, can be obtained from pooled plasma. Fibrin scaffolds form when thrombin cleaves fibrinogen into fibrin monomers that assemble to form a non-covalent scaffold. Factor XIIIa then creates covalent cross-links, which stabilize the scaffold. Similar to collagen, fibrin scaffolds contain sites for cell adhesion and the scaffold properties vary depending on the concentration of fibrin used. Additionally, fibrin can also be covalently modified to further alter its properties (Willerth and SakiyamaElbert 2007). Most of the time, the latter is bound to growth factors that can elecit axonal regeneration. One of the most common methods to load growth factors into the fibrin hydrogels is the one previously described by SakyiamaElbert and Hubbel (Taylor et al., 2004) who have developed an affinitybased drug delivery system designed to slow the diffusion of heparin-binding growth factors from fibrin gels. The system allows the release of growth factors to be controlled by cell-mediated processes, such as cell-activated plasmin degradation. The heparin-based delivery system (HBDS) has three components: a synthetic linker peptide, the polysulphated glycosaminoglycan heparin, and the growth factor to be delivered. An example of this strategy was presented by Taylor et al. (2004). The

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authors assessed the effect of controlled delivery of NT-3 from fibrin scaffolds implanted in spinal cord lesions immediately following a 2-mm ablation injury in adult rats. Results revealed that nine days after injury, fibrin scaffolds containing the delivery system elicited more robust neuronal fibre growth into the lesion than did control scaffolds or saline. Implantation of fibrin scaffolds resulted in a dramatic reduction of glial scar formation at the white matter border of the lesion.

22.5

Conclusions and future trends

The CNS, and within it the spinal cord, are extremely complex. Therefore only an integrative multidisciplinary approach can deal with the complexity of its regeneration. As was described, several attempts have been made, but none up to now has made it into the clinics. The most probable cause for this is the fact that most of them were based on the application of one strategy (cell based therapies, pharmacological agents, hydrogel based biomaterials) instead of a combination of all of them. Therefore a special effort needs to be made in order to combine all these disciplines. This has already been done with success in other tissues. A better characterization of adult stem cell and OECs must also be further addressed. Knowledge of the mechanisms regulating cell differentiation and neuronal pathways is still poor and new techniques are needed for their characterization, purification and expansion. New sources of cells are needed in order to overcome some of the limitations of the NSCs isolated from the CNS. The second aspect that needs to be improved is related to material science. A new generation of biodegradable biomaterials is currently being designed, to elicit specific cellular responses at the molecular level. This third generation of biomaterials is based on molecular modifications of biodegradable polymer systems that will later stimulate specific interactions with cell integrins and thereby direct cell proliferation, differentiation and extracellular matrix production and organization. Self-assembled materials are also another class of materials that can be used for tissue engineering purposes. Particularly interesting is the work revealed by Stupp’s laboratory (Silva et al., 2004). Neural progenitor cells were encapsulated in vitro within a three-dimensional network of nanofibres formed by self-assembly of peptide amphiphile molecules. These nanofibres were designed to present to cells the neuritepromoting laminin epitope IKVAV. Relative to laminin or soluble peptides, the artificial nanofibre scaffold induced very rapid differentiation of cells into neurons, while inhibiting astrocyte differentiation. This rapid selective differentiation is linked to the amplification of bioactive epitope presentation to cells by the nanofibres. In conclusion it can be said that hydrogel based biomaterials have in fact

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a tremendous potential to overcome the shortcomings of the existing therapies for SCI regeneration. However, the next years will be decisive for their affirmation within the SCI field. Because of this it is necessary to improve the interaction of the latter with more fundamental aspects of SCI and the cells associated to its possible regeneration so that the final goal, the production of SC equivalents that lead to SCI regeneration, can be achieved.

22.6

Acknowledgements

Portuguese Foundation for Science and Technology through funds from POCTI and/or FEDER programs (post doctoral fellowship to A J Salgado - SFRH/ BPD/17595/2004). The authors also acknowledge Luis Osório for the spinal cord illustrations.

22.7

References

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lesion in adult rats: dose-window and time-window effects’, J Neurotrauma 17(1), 19–30. Grothe C, Meisinger C and Claus P (2001), ‘In vivo expression and localization of the fibroblast growth factor system in the intact and lesioned rat peripheral nerve and spinal ganglia’, J Comp Neurol, 434, 342–357. Grothe C, Haastert K and Jungnickel J (2006), ‘Physiological function and putative therapeutic impact of the FGF–2 system in peripheral nerve regeneration. Lessons from in vivo studies in mice and rats’, Brain Res Brain Res Rev, 51(2), 293–299. Guest J D, Hiester E D and Bunge R P (2005), ‘Demyelination and Schwann cell responses adjacent to injury epicenter cavities following chronic human spinal cord injury’, Exp Neurol, 192, 384–393. Guth L, Barrett C P, Donati E J, Deshpande S S and Albuquerque E X (1981), ‘Histopathological reactions and axonal regeneration in the transected spinal cord of hibernating squirrels’, J Comp Neurol, 103, 297–230. Hall E D and Wolf D L (1986), ‘A pharmacological analysis of the pathophysiological mechanisms of posttraumatic spinal cord ischemia’, J Neurosurg, 64(6), 951–961. Hearn M T W (1987), 1, 1-carbonyldiimidazole-mediated immobilization of enzymes and affinity ligands, Methods Enzymol, 135, 102–117. Holaday J W and Faden A I (1980), ‘Naloxone acts at central opiate receptors to reverse hypotension, hypothermia and hypoventilation in spinal shock’, Brain Res, 189(1), 295–300. Houle J D and Tessler A (2003), ‘Repair of chronic spinal cord injury’, Exp Neurol, (2003), 182, 247–260. Iannotti C, Li H, Yan P, Lu X, Wirthlin L and Xu X M (2003), ‘Glial cell line-derived neurotrophic factor-enriched bridging transplants promote propriospinal axonal regeneration and enhance myelination after spinal cord injury’, Exp Neurol, 183(2), 379–393. Ishii K, Nakamura M, Dai H, Finn T P, Okano H, Toyama Y and Bregman B S (2006), ‘Neutralization of ciliary neurotrophic factor reduces astrocyte production from transplanted neural stem cells and promotes regeneration of corticospinal tract fibers in spinal cord injury’, J Neurosci Res, 84(8), 1669–1681. Iwanami A, Kaneko S, Nakamura M, Kanemura Y, Mori H, Kobayashi S, Yamasaki M, Momoshima S, Ishii H, Ando K, Tanioka Y, Tamaoki N, Nomura T, Toyama Y and Okano H (2005), ‘Transplantation of human neural stem cells for spinal cord injury in primates’, J Neurosci Res, 80, 182–190. Jasmin L, Janni G, Moallem T M, Lappi D A and Ohara P T (2000), ‘Schwann cells are removed from the spinal cord after effecting recovery from paraplegia’, J Neurosci, 20, 9215–9223. Jimenez Hamann M C, Tator C H and Shoichet M S (2005), ‘Injectable intrathecal delivery system for localized administration of EGF and FGF-2 to the injured rat spinal cord’, Exp Neurol, 194(1), 106–119. Junqueira L and Carneiro J (2004), Basic Histology, 10th Edition, Guanabana Koogan, Rio de Janeiro. King V R, Henseler M, Brown R A and Priestley J V (2003), ‘Mats made from fibronectin support oriented growth of axons in the damaged spinal cord of the adult rat’, Exp Neurol, 182(2), 383–398. King V R, Phillips J B, Hunt-Grubbe H, Brown R and Priestley J V (2006), ‘Characterization of non-neuronal elements within fibronectin mats implanted into the damaged adult rat spinal cord’, Biomaterials, 27(3), 485–496.

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Knoblach S M, Huang X, VanGelderen J, Calva-Cerqueira D and Faden A I (2005), ‘Selective caspase activation may contribute to neurological dysfunction after experimental spinal cord trauma’, J Neurosci Res, 80, 369–380. Kohama I, Lankford K L, Preiningerova J, White F A, Vollmer T L and Kocsis J D (2001), ‘Transplantation of cryopreserved adult human Schwann cells enhances axonal conduction in demyelinated spinal cord’, J Neurosci, 21, 944–950. Kohn J and Wilchek M (1981), ‘Procedures for the analysis of cyanogen bromide-activated Sepharose or Sephadex by quantitative determination of cyanate esters and imidocarbonates’, Anal Biochem, 115, 375–382. Koshizuka S, Okada S, Okawa A, Koda M, Murasawa M, Hashimoto M, Kamada T, Yoshinaga K, Murakami M, Moriya H and Yamazaki M (2004), ‘Transplanted hematopoietic stem cells from bone marrow differentiate into neural lineage cells and promote functional recovery after spinal cord injury in mice’, J Neuropathol Exp Neurol, 63(1), 64–72. Kuo H S, Tsai M J, Huang M C, Huang W C, Lee M J, Kuo W C, You L H, Szeto K C, Tsai I L, Chang W C, Chiu C W, Ma H, Chak K F and Cheng H (2007), ‘The combination of peripheral nerve grafts and acidic fibroblast growth factor enhances arginase I and polyamine spermine expression in transected rat spinal cords’, Biochem Biophys Res Commun, 357(1), 1–7. Kuroki K, Cook J L, Kreeger J M and Tomlinson J L (2003), Osteoarthritis Cartilage, 11, 625–635. Kwon B K, Tetzlaff W, Grauer J N, Beiner J and Vaccaro A R (2004), ‘Pathophysiology and pharmacologic treatment of acute spinal cord injury’, The Spine Journal, 4, 451– 464. Luo J, Borgens R and Shi R (2002), ‘Polyethylene glycol immediately repairs neuronal membranes and inhibits free radical production after acute spinal cord injury’, J Neurochem, 83, 471–480. Luo J and Shi R (2004), ‘Diffusive oxidative stress following acute spinalcord injury in guinea pigs and its inhibition by polyethyleneglycol’, Neurosci Lett, 359, 167–170. Luo Y and Shoichet M S (2004a), ‘Light-activated immobilization of biomolecules to agarose hydrogels for controlled cellular response’, Biomacromolecules, 5(6), 2315– 2323. Luo Y and Shoichet M S (2004b), ‘A photolabile hydrogel for guided three-dimensional cell growth and migration’, Nat Mater, 3(4), 249–253. Luo J Borgens R and Shi R (2004), ‘Polyethylene glycol improves function and reduces oxidative stress in synaptosomal preparations following spinal cord injury’, J Neurotrauma, 21, 994–1007. Mader S (2004), Understanding Human Anatomy and Physiology, 5th edn, Mcgraw Hill Companies. Mahoney M and Anseth K S (2007), ‘Contrasting effects of collagen and bFGF-2 on neural cell function in degradable synthetic PEG hydrogels’, J Biomed Mat Res, 81A, 268–278. Marieb E (1992), Human Anatomy and Physiology, Redwood City, CA: Benjamin/ Cummings, (pp. 491–492). Marieb E (2004), Human Anatomy & Physiology (6th ed.), New York: Pearson. Merkle F T, Mirzadeh Z and Alvarez-Buylla (2007), ‘Mosaic organization of neural stem cells in the adult brain’, Science, 317(5836), 381–384. Miron T and Wilchek M (1987), ‘Immobilization of Proteins and ligands using chlorocarbonates’, Methods Enzymol, 135, 84–90.

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Miyoshi Y, Date I, Ohmoto T and Iwata H (1996), ‘Histological analysis of microencapsulated dopamine-secreting cells in agarose/poly(styrene sulfonic acid) mixed gel xenotransplanted into the brain’, Exp Neurol, 138, 169–175. Netter F H (2003), Atlas of Human Anatomy, 3rd Edn, Teterboro, NJ: Icon Learning Systems Nguyen K T and West J L (2002), ‘Photopolymerizable hydrogels for tissue engineering applications’, Biomaterials, 23(22), 4307–4314. Novikova L N, Mosahebi A, Wiberg M, Terenghi G, Kellerth J O and Novikov L N (2006), ‘Alginate hydrogel and matrigel as potential cell carriers for neurotransplantation’, J Biomed Mater Res A, 77(2), 242–252. Ogawa Y, Sawamoto K, Miyata T, Miyao S, Watanabe M, Nakamura M, Bregman B S, Koike M, Uchiyama Y, Toyama Y and Okano H (2002), ‘Transplantation of in vitro-expanded fetal neural progenitor cells results in neurogenesis and functional recovery after spinal cord contusion injury in adult rats’, J Neurosci Res, 69, 925– 933. Orive G, Ponce S, Hernandez R M, Gascon A R, Igartua M and Pedraz J L (2002), ‘Biocompatibility of microcapsules for cell immobilization elaborated with different type of alginates’, Biomaterials, 23, 3825–3831. Peppas N A, Keys K B, Torres-Lugo M and Lowman A M (2002), ‘Poly(ethylene glycol)containing hydrogels in drug delivery’, J Control Release, 62(1-2), 81–87. Phillips J B, King V R, Ward Z, Porter R A, Priestley J V and Brown R A (2004), ‘Fluid shear in viscous fibronectin gels allows aggregation of fibrous materials for CNS tissue engineering’, Biomaterials, 25, 2769–2779. Piantino J, Burdick J A, Goldberg D, Langer R and Benowitz L I (2006), ‘An injectable, biodegradable hydrogel for trophic factor delivery enhances axonal rewiring and improves performance after spinal cord injury’, Exp Neurol, 201(2), 359–367. Pinzon A, Calancie B, Oudega M and Noga B R (2001), ‘Conduction of impulses by axons regenerated in a Schwann cell graft in the transected adult rat thoracic spinal cord’, J Neurosci Res, 64, 533–541. Plato (1993), The Republic, Trans MHR Pereira, Lisboa: Fundação Calouste Gulbenkian. Prang P, Müller R, Eljaouhari A, Heckmann K, Kunz W, Weber T, Faber C, Vroemen M, Bogdahn U and Weidner N (2006), ‘The promotion of oriented axonal regrowth in the injured spinal cord by alginate-based anisotropic capillary hydrogels’, Biomaterials, 27(19), 3560–3569. Purves D (1997), Neuroscience, Sinauer Associates Publishing, New York. Ramón-Cueto A and Nieto-Sampedro M (1994), ‘Regeneration into the spinal cord of transected dorsal root axons is promoted by ensheathing glia transplants’, Exp Neurol, 127(2), 232–244. Ramón-Cueto A, Cordero M I, Santos-Benito F F and Avila J (2000), ‘Functional recovery of paraplegic rats and motor axon regeneration in their spinal cords by olfactory ensheathing glia’, Neuron, 25(2), 425–435. Rasouli A, Bhatia N, Suryadevara S, Cahill K and Gupta R (2006), ‘Transplantation of preconditioned Schwann cells in peripheral nerve grafts after contusion in the adult spinal cord. Improvement of recovery in a rat model’, J Bone Joint Surg Am, 88(11), 2400–2410. Ray S K, Matzelle D D, Sribnick E A, Guyton M K, Wingrave J M and Banik N L (2003), ‘Calpain inhibitor prevented apoptosis and maintained transcription of proteolipid protein and myelin basic protein genes in rat spinal cord injury’, J Chem Neuroanat, 26, 119–124.

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synthetic hydrogel guidance channels improves specific supraspinal and local axonal regeneration after complete spinal cord transection’, Biomaterials, 27(3), 519–533. Van De Graaff K M (2001), Human Anatomy, 6th Edition, McGraw-Hill Companies. Vergara C and Ramirez B (20004), ‘CNTF, a pleiotropic cytokine: emphasis on its myotrophic role’, Brain Res Brain Res Rev, 47(1–3), 161–173. Wang B, Zhao Y, Lin H, Chen B, Zhang J, Zhang J, Wang X, Zhao W and Dai J (2006), ‘Phenotypical analysis of adult rat olfactory ensheathing cells on 3-D collagen scaffolds’, Neurosci Lett, 401(1-2), 65–70. Willerth S M and Sakiyama-Elbert S E (2007), ‘Approaches to neural tissue engineering using scaffolds for drug delivery’, Adv Drug Delivery Rev, 59, 325–338. Woerly S (2000), ‘Restorative surgery of the central nervous system by means of tissue engineering using NeuroGel implants’, Neurosurg Rev, 23(2), 59–77. Woerly S, Pinet E, de Robertis L, Bousmina M, Laroche G, Roitback T, Vargová L and Syková E (1998), ‘Heterogeneous PHPMA hydrogels for tissue repair and axonal regeneration in the injured spinal cord’, J Biomater Sci Polym Ed, 9(7), 681–711. Woerly S, Petrov P, Syková E, Roitbak T, Simonová Z and Harvey A R (1999), ‘Neural tissue formation within porous hydrogels implanted in brain and spinal cord lesions: ultrastructural, immunohistochemical, and diffusion studies’, Tissue Eng, 5(5), 467– 488. Woerly S, Pinet E, de Robertis L, Van Diep D and Bousmina M (2001a), ‘Spinal cord repair with PHPMA hydrogel containing RGD peptides (NeuroGel)’, Biomaterials, 22, 1095–1111. Woerly S, van Diep D, Sosa N, de Vellis J and Espinosa A (2001b), ‘Reconstruction of the transected cat spinal following NeuroGelTM’, Int J Develop Neurosci, 19, 63–83. Xiao M, Klueber K M, Lu C, Guo Z, Marshall C T, Wang H and Roisen F J (2005), ‘Human adult olfactory neural progenitors rescue axotomized rodent rubrospinal neurons and promote functional recovery’, Exp Neurol, 194, 12–30. Yang Y B and Piao Y J (2003), ‘Effects of resveratrol on secondary damages after acute spinal cord injury in rats’, Acta Pharmacol Sin, 24, 703–710. Yoshii S, Oka M, Shima M, Taniguchi A, Taki Y and Akagi M (2004), ‘Restoration of function after spinal cord transection using a collagen bridge’, J Biomed Mater Res A, 70(4), 569–575. Young W, Flamm E S, Demopoulos H B, Tomasula J J and DeCrescito V (1981), ‘Effect of naloxone on posttraumatic ischemia in experimental spinal contusion’, J Neurosurg, 552, 209–219. Young W (2000), ‘Molecular and cellular mechanisms of spinal cord injury therapies’, in Neurobiology of spinal cord injury, Kalb R and Strittmatter S M, Humana Press, Totowa, NJ, 241–276. Zhang L, Gu S, Zhao C and Wen T (2007), ‘Combined treatment of neurotrophin-3 gene and neural stem cells is propitious to functional recovery after spinal cord injury’, Cell Transplant, 16(5), 475–481.

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Part V Systems for the sustained release of molecules

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23 Particles for controlled drug delivery E. T. B A R A N and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

23.1

Introduction

In many diseases, drugs must be administered at levels in the range between effective and toxic to avoid ineffective or toxic dosages. In some diseases, such as diabetes or thyroid disease, drugs must also be administered at specific time intervals. These factors lie behind the need for a smart delivery system that can respond to changes in the patient’s metabolism and release the drug accordingly. The encapsulation of drugs in biodegradable polymers is considered a promising method for controlled delivery. This drug particulate delivery method has an additional advantage in the potential for targeting delivery by modifying the particle surfaces with biological ligands or high affinity molecules, such as antigens or antibodies. Prolonged continuous drug release can be obtained using spherical microparticles, due to their high volume to surface ratio. These microparticle drug carriers can be made from a diverse variety of natural and synthetic biodegradable polymers (Table 23.1). Nanoparticles with unique properties can also be prepared using biodegradable polymers. The naturally occurring polymers used in nanoparticle preparation and drug release are shown in Table 23.2. Nanoparticles are more suitable than microparticles for intravenous administration and targeting drugs by systemic transport. The delivery of relatively stable, small molecular weight drugs and unstable macromolecular bioactive agents, such as growth hormone, interferons, and plasmids, requires the use of suitable micro encapsulation techniques and polymeric micro carriers. The efficient delivery of drugs, particularly to diseased body sites, is a major challenge in drug delivery as the complex body environment and hostile immune system effectively prevent micro carriers from reaching their target site.

23.2

Novel particle processing methods

Microparticles have been investigated extensively as a new drug delivery method. By encapsulating a drug in a biodegradable, microparticular form 597 © 2008, Woodhead Publishing Limited

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Table 23.1 Biodegradable polymers used in microparticle preparation for drug delivery Polymer and macromolecules Polysaccharides Alginate Chitosan Carboxymethyl-Chitin Carboxymethyl-Chitosan Propyl, Butyl chitosan Chondroitin Dextran Ethylcellulose Hyaluronan Pectin Pullunan Starch Proteins Albumin Avidin Casein Collagen, gelatin Corn protein Silk fibroin Ovalbumin Whey protein Synthetic polymers Poly(capro lactone) Polyanhydrides Polycarbonate Poly(lactic acid) Poly(lactic glycolic acid)

Poly(organophosphozanes) Poly(phosphoester) Poly(urethanes) Lipids and amphoteric molecules Lauryl dextrin Phosphotidylcholine, soya oil Poly(palmitoyl-Lhydroxyproline ester) Stearic acid

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Encapsulated drug and reference(s) Isoniazid,1 chymotrypsin,2 tetramidine,3 DNA4 Ketoprofen,5 felodipine,6 vaccine,7 6-Mercaptopurine8 Pazufloxacin,9 BSA,10 Mercaptourine11 Metaclopramide12 Theophyline,13 bone morphpgenic protein,14 mitomycinC,15 (interleukin-2)16 Diclofenac,17 DNA,18 clomolyn,19 gentamicin20 Piroxicam,21 sulphanilamide-sulphaguanidinesulphathiozole22 β-Estradiol-propranolol-metronidazol-diclofenacindomethacin23 Cisplatin,24 contrast agent25 Bupivacaine,26 mitoxantrone,27 clarithromycin28 Glycoprotein29 Progesteron30 Deoxyuridine,31 insulin,32 fibroblast growth factor33 Ivermectin34 Horseradish peroxidase35 Muramyl dipeptide36 Folate37 Papaverine,38 bupivacaine,39 deoxycycline,40 ketoprofen,41 nifedipine-proprandol42 p-Nitrophenol-lysozyme,43 BSA,44 p-nitroaniline45 Piroxicam,46 N-Cyclopentyladenosine,47 holmium acetylacetonate,48 nimesulide49 Proxicam,50 gentamicin51 temozolomide,52 5-flurouracil,53 doxorubicin,54 insulin,55 degarelix56 Insulin,57 naproxen58 Nerve growth factor59 Diazinon,60 benzocaine61

Cyclosporin62 ProstaglandinE1,63 indomethacin64 Rifampicin65 Cefuroxime axetil66

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Table 23.2 Natural biodegradable polymers used in nanoparticle preparation for drug delivery Polymer and macromolecules

Encapsulated bioactive agent and Reference(s)

Polysaccharides Alginate Chitosan Pectinate Wheat gliadin

DNA,67 insulin68 Insulin,69 oligonucleotides and DNA,70 doxorubicin71 Insulin72 α-Tocopherol73

Proteins Albumin Ferritin Gelatin Hydrophobicaly modified poly(γ-glutamic acid) Silk sericin Vicilin storage protein

Antisense oligonucleotides,74 Quantum dot75 Oligonucleotides,76 paclitaxel77 Ovalbumin-recombinant human immunodefiviency virus78 –79 –80

Lipids Chlosterol-pullulan Preciol Tristearin glyceride Bacterial polymers Poly(3-hydroxybutyrateco-valerate)

Insulin81 Ubiprofen82 Triptolide83 Asparaginase-catalase84

(microsphere, microcapsule), a sustained drug release profile can be maintained. Much of the research to date has been on obtaining microspheres and microcapsules using biodegradable and biocompatible polymers. Drug release can be modulated through polymer chemistry and biodegradability, and also the physical shape and processing conditions (the method of preparation, crosslinking, freeze-drying, etc.). However, the ability to tailor drug release kinetics solely through polymer chemistry and processing conditions is limited. The size of the microsphere is a primary factor governing drug release under identical conditions of porosity and polymer molecular weight. Larger microspheres generally release drugs at a slower rate over longer time periods.85 Microparticle preparation techniques that can produce drug carriers in narrow and adjustable size distributions can be used to control drug release rates. Traditional micro-encapsulation techniques, such as emulsion/solvent evaporation, spray drying and particle milling techniques, often use high temperatures, toxic solvents and mechanical stresses that cause degradation of the drug and high particle size variations. For these reasons, there is great interest in new microencapsulation techniques that control particle size but use smaller mechanical forces, less energy input and little or no organic solvent.

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In order to achieve a narrow particle size distribution, engineered surface properties, and environmentally acceptable processing conditions, researchers have turned to new technologies and techniques that have been developed for micro-fabrication of electronic devices, including laser processing, and simple but effective techniques such as enveloping particles with thin polymer films. In particular, recent developments in micro-fabrication methods and nanotechnology provide precise tools for the production of microparticles with engineered surfaces and bulk properties. Recently, biodegradable mono-dispersed polymer particles ranging in size from several tens of microns to nanometres were made by electrodynamic atomization, to be used for the sustained delivery of the anticancer drug paclitaxel to treat C6 glioma in vitro.86 Electrospraying or electrodynamic atomization is based on a potential difference created between the nozzle and a high voltage ring. A solution containing dissolved polymer and the drug emitted from the nozzle forms a liquid cone with a thin jet, which breaks up into mono-dispersed droplets. Similarly, a piezoelectric array microjet, in which a fluid chamber is formed by a piezoelectric actuator bonded to a silicon chip with nozzles, can be used to produce microparticles within a narrow size range for drug delivery applications.87 According to the principle of droplet delivery used in inkjets, a pressure wave generated by the piezoelectric actuator inside the fluid chamber propagates towards the nozzles and squeezes the liquid out of the chamber. With sufficient kinetic energy, the liquid is then released through the nozzles and forms homogeneous polymer droplets. Nano-functionalized drug particles have been produced using laser pulsed ablation to control the architecture and thickness of a nanoscale polymer coating on the drug particles.88 Drug particles (budesonide, 1-5 µm size) with a nanoscale, biodegradable polymer coating (PLGA) were held in a fluidized bed so that their entire surface was exposed to the incoming flux. In vitro dissolution tests showed that a significant slow down of release was observed with coated particles compared with uncoated drug particles. A microporous membrane with controlled pore size can distribute a specific size of emulsion droplets at lower energy input and without high mechanical stress.89 The emulsions are produced by pressurizing the dispersed phase into a continuous phase through a microporous membrane, and the emulsion droplet size is controlled by the membrane pore size. This membrane emulsification was used to make a water in oil in water (w/o/w) emulsion of poppy seed oil (lipiadol), to encapsulate the anticancer drug epirubicin for the treatment of liver cancer.90 The drug emulsion was injected into the patient’s hepatic artery, and subsequent selective deposition in the cancerous tumour was demonstrated. A similar but more advanced technique, microchannel emulsification, has been used to produce mono-dispersed polymeric microspheres.91 This technique

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is based on micro-fabricated microchannel arrays created using photolithographical methods. Microchannel processing was used to manufacture mono-dispersed chitosan microparticles containing amplicilin.92 In this study, a chitosan aqueous phase (the disperse phase) was fed into a microfluidic chip equipped with a cross-junction microchannel, and sheared by viscous oil flow (the continuous phase). The microparticles produced ranged in size from 100 to 800 µm with a variation of less than 5%, and predictable release patterns as a function of particle size were obtained. Production of micron-sized mono-dispersed poly-(lactide-co-glycolide) (PLGA) and poly(l-lactic acid) terminated by 1H, 1H-perfluoro octyloxycarbonyl (PLA-PFO) microspheres using ink-jet printing technology was studied.93 Droplets of polymer solutions in volatile solvent were printed with a piezo-driven ink-jet nozzle which was submerged in an aqueous phase, and a constant pressure in the nozzle was applied for a high droplet formation rate. The emulsion droplets were collected at the bottom of the aqueous phase and post-treated to wash away the emulsifier and allow the solvent to evaporate. The potential of supercritical processes for drug delivery applications has also attracted attention recently. Control over particle size, absence of toxic solvents and preservation of drug activity and polymer molecular weight are the main advantages supercritical methods have over conventional production of drugs encapsulated in polymeric carriers. In particular, supercritical CO2 (scCO2) has a number of unique properties that make it possible to process both bioactive molecules and amorphous polymers without using elevated temperatures. In one method, the drug and the polymer are dissolved in scCO2 in a pressurized chamber and subsequently the solution is sprayed through a nozzle to produce fine drug-polymer particles in a depressurized collection chamber.94 Alternatively, scCO2 can be used as an anti-solvent for the precipitation of drugs already dissolved in organic solvents. Sustained release behaviour was obtained by spraying hydrocortisone particles along with PLGA solution (in dichloromethane) into scCO2 as an antisolvent, which resulted in the drug particles being coated with a thin polymer layer (<20 µm).95 Recently, the bacterial polymer poly(3-hydroxyco-3-hydroxyvalerate) in dichloromethane solution was micronized and precipitated using the supercritical antisolvent technique to produce spherical microparticles of 3–9 µm in diameter.96 As a hydrophilic biopolymer, aqueous chitosan solution has been micronized using supercritical assisted atomization to produce microparticles.97 The particles were found to be spherical and ranged in size between 0.1 to 1.5 µm when precipitated using the supercritical fluid process. Microfabrication techniques that permit the creation of therapeutic delivery systems with a combination of structural, mechanical and perhaps electronic features may surmount the challenges presented by conventional delivery of

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therapy.98 This technology can be used to fabricate particulate drug delivery microdevices with highly uniform sizes, and well-defined and asymmetrical structures that are impossible to create by conventional microparticle processing methods. The particles can be asymmetrically designed with single or multiple drug reservoirs, and ligands bound to one side can be used for targeting the desired delivery site. Micro-fabricated microparticles made up of poly(methyl methacrylate) in thin square structures (150 × 150 × 5 µm) with reservoir (80 × 80 × 2 µm) for drug encapsulation were reported by Tao et al.99 Photolithography methods with ion etching techniques were used to create the particles. In a second photolithography step, the reservoirs were built inside microparticle boundaries and subsequently they were surface modified with lectin to target the particles to the CaCo-2 cell line as a model gastrointestinal track. Another micro-fabrication method, soft lithography, uses an elastomeric stamp with topographical features to transmit micro/nano structures, and this can be applied to the processing of biodegradable polymers for the production of microparticles. By combining microcontact printing, microtransfer moulding, and lift-off techniques, polymeric microparticles with engineered features have been produced.100 Using these techniques, PLGA, chitosan, and poly(ethylene glycol dimethacrylate) microparticles with a wide range of sizes (5–198 µm) and shapes have been fabricated.

23.3

Hiding particles: The stealth principle

Particulate drug formulations are delivered via several routes, such as oral,101 ocular,102 intra-nasal,103 dermal,104 parenteral,105 and pulmonary.106 For deep tissue delivery, parenteral administration of drugs and carriers has become more important as the carriers can be directed systemically to the target organ, tissue, or even particular cell types. Because of the complex biochemical environment in the body, and the effective clearance of foreign molecules by the reticuloendothelial (RES) system, however, parenteral delivery of particulate drug carriers is highly challenging. Stealth-targeting for cellular delivery is essential.

23.3.1 Stealth particles and the biomimetic approach for increasing circulation half-life Strategies to avoid elimination of drugs by the immunological defence mechanism include coating the particles with hydrophilic polymers such as PEG, which act as steric barriers, and biomimetic approaches, such as coating the particles in a phospholipid bilayer which resembles that of cell membranes, artificial viruses, and virus-like particles. Foreign particles in the circulatory system are quickly recognized and

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removed by mononuclear leukocytes (the mononuclear phagocyte system and polymorphonuclear leukocytes). Macrophages located in the RES system play an essential role in removing injected particles, with an efficiency of 85–95%.107 Blood components (opsonins) on the surface of particles, for example serum complement proteins, in particular C3, strongly activate phagocytosis by macrophages.108 It is, therefore, essential to modify the surface of the drug delivery particles with macromolecules that can repel the adsorption of proteins from plasma. The best-known hydrophilic, neutral polymer that prevents protein adsorption effectively is poly(ethyleneglycol) (PEG), which shows great potential for increasing the circulation half-life of drug delivery particles.109–111 Studies with liposomes and block co-polymers which have a hydrophobic stable core and hydrophilic block that forms a repulsive and hydrophilic layer, suggest that the flexibility of the protective polymer may be important for the effectiveness of PEG. Liposomal doxorubicin Caelyx/Doxil® was studied for clinical use in the treatment of refractory Kaposi’s sarcoma and ovarian cancer.112 Caelyx® consists of doxorubicin hydrochloride encapsulated in a stealth liposome, which contains the phospholipid, hydrogenated soybean phosphatidylcholine. The surface of the liposome is coated with polymers of PEO. Caelyx® showed stealth-like properties, which result in prolonged plasma circulation and an expected reduction in doxorubicin-related toxicities. Based on phase II studies in ovarian cancer, Caelyx® seems to offer a reduced toxicity profile. The clinical data113 suggested that Caelyx® had a significant level of activity in the difficult refractory population, with an overall response rate between 9 and 15% in those who relapsed within six months. It was also reported to be well tolerated, with minimal evidence of myelosuppression, alopecia or cardiac toxicity.

23.3.2 The biomimetic approach in surface modification of particles Biological functions, such as material exchange across the cell membrane, combined with recent achievements in membrane science, are providing important cues to investigators seeking to build more biologically friendly drug carriers. The red blood cell (RBC) has become a useful model, as it has a surface that is inert to serum protein adsorption and is therefore immune to capture by reticuloendothelial cells. Generally, RBC carriers for microspheres are prepared by collecting a blood sample from the organism of interest, separating the erythrocytes from plasma, removing the leukocytes, enclosing the drug in the erythrocytes, and finally resealing the resulting cellular carriers.114 In vitro studies with canine blood containing adriomycine obtained by hypotonic dialysis showed a slow release of the drug from erythrocytes treated with gluteraldehyde for

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crosslinking.115 Red blood cells can be used as carriers for the anti-leukemic drug L-asparaginase, which can otherwise provoke severe immunogenic reactions in patients upon repeated administration because of its protein nature.116 Red blood cells carrying L-asparaginases were detected to have a circulation half-life similar to native RBCs. The targeting potential of RBCs after surface modification was studied by Mishra et al.117 Erythrocytes encapsulated with methotrexate were surface modified using desialation and hemichrome induction by treating with trypsin (Tt) and phenylhydrazine (PhT), respectively. The macrophage uptake of both surface-modified erythrocytes was reported to be enhanced by 3 to 5 and 5 to 6 times with Tt and PhT cells, respectively. The possible use of RBCs as a transient transport vehicle to deliver drugs to distant targets for which hook peptides have higher affinity was studied by Feder et al.118 The binding of antimicrobial peptide, a derivative of antimalarial microparticle dermaceptin S4, K4-S4 (1–13)a, to the plasma membrane of RBCs was assessed in vitro and in vivo. The transfer of the RBC-bound peptide to the plasma membrane of microorganisms was reported to be rapid, spontaneous and internalized, with a receptor independent mechanism. Phospholipid assemblies on particular drug carriers can mimic biological membranes. A biovector biomimetic synthetic delivery system, known as the light supramolecular biovector (SMBV), is a successful example of this approach.119 These nanoparticles mimic viruses, such as influenza, in terms of their supramolecular structure, with a lipid bilayer surrounding an internal polysaccharide core where drugs can be encapsulated.120 Lipid-coated, biomimetic hydrogel particles can have unique release properties. A crosslinked polyanionic polymeric microparticle encapsulated within a lipid membrane was investigated by Kiser et al.121 The lipid-coated hydrogel microparticles were triggered to expose and exchange their encapsulated doxorubicin with external fluid over a period of several minutes by high voltage electroporation of the phospholipid bilayer.

23.4

Finding the target

Several approaches can be used to ensure that drug-loaded particles reach the target tissue or particular cell types (the magic bullet principle) and safely deliver the drug. For example, the ligand molecules specific to certain cell receptors and the antibodies directed against antigens on the cell surface can be utilized to modify the particle surface and promote bio-affinity interactions. This active targeting strategy uses the exceptional ability of biologically active molecules to bind to specific, complementary active sites. Likewise, by rendering the particles vulnerable to external forces, such as magnetic fields and radiation, the drug carriers can be localized and even triggered to release the drug at the target site. Alternatively, the size of the

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particles can be optimized so that they accumulate at certain body sites after injection. Passive targeting is based on the tendency of microparticles to entangle in tissue structures such as capillary beds, tumour tissues or lung alveoli. The tendency of microparticles to localize in RES in liver and spleen is an example of passive targeting. The particulate carriers accumulate in the tumour due to its leaky vasculature and poorly developed lymphatic system, a result known as enhanced penetration and retention (EPR).122 For instance, PEG-stabilized poly(propylene sulphide) nanoparticles (20-100 nm), designed to deliver antigens to antigen-presenting cells in lymph nodes, were detected after interstitial injection to be at an internalization efficiency of between 40-50% without the use of a targeting ligand.123 The lymphatic system is involved in the dissemination of cancer cells. For this reason, an effective trans-lymphatic drug delivery system using controlled release formulations may enhance drug targeting to cancer cells while reducing systemic toxicity. Using a passive targeting strategy, a translymphatic drug delivery system that incorporated PLGA-paclitaxel or PLGA-rhodamine microspheres in a gelatine sponge matrix showed spontaneous absorption of the particles in the lymphatic systems of rats bearing orthotopic lung cancer.124 Similarly, lipid microspheres (soy bean oil with lecithin) encapsulating 1,3bis (2-chloroethyl)-1-nitrosourea (BCNU) showed significant anti-tumour activity in mice with 1210 leukaemia.125 Smaller nanospheres (50 nm) showed a similar anti-tumour activity and lower uptake by reticuloendothelial cells, indicating lower non-specific delivery.

23.4.1 Magnetic targeting Magnetic nanoparticles can bind to drugs, proteins, enzymes, antibodies, or DNA, and can be directed to specific organ sites, tissues, or tumours using an external magnetic field.126 A hypothetical drug targeting system, utilizing high gradient magnetic separation principles was studied by Ritter et al.127 The results revealed that a magnetic targeting approach using ferromagnetic wire to magnetically collect particles at a target site (even in arteries with very high velocities) is feasible, but for magnetically guiding a particle through the circulatory system there are limitations, especially for magnetic systems applied externally. In another study, a mathematical model was used to examine the capture of magnetic drug carrier particles by a magnetizable intravascular stent (MIS) on parameters such as blood flow rate, magnetic field strength, particle properties and stent design.128 The results suggested that this tool could be effective in bringing drugs to various disease sites. Grief et al. formulated a model of magnetic particle transport in the intermediate sized vessels of the blood stream that incorporates the effects of shear-induced diffusion (which arises as a result of interactions between red

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blood cells).129 Their model predicted that the use of magnetically targeted delivery with an externally applied magnetic field is more appropriate for targets close to the body surface, as the magnitude of the magnetic field required to hold a particle in the main flow of all but the very smallest of vessels is very large. In a similar approach, the feasibility of magnetically targeting the deposition of aerosols for application in lung cancer was investigated by Ally et al.130 According to the study, the particle concentrations are important and the high-gradient region of the magnetic field strength and the gradient should be as large as possible, with the in vitro model demonstrating 48–87% collection efficiency. Superparamagnetic, iron-based molecules can be used for magnetic carriers. Alginate microspheres with trapped iron oxide nanoparticles and iron ions bound to the polymer structure showed paramagnetic properties suitable for magnetic targeting.131 Similarly, intra-arterially infused doxorubicin-loaded carbon-metal iron particles can be localized and retained within a tumour mass by an externally positioned, permanent dipole magnet.132 Clinical studies of the treatment of hepatocellular carcinoma with magnetic particles demonstrated good particle tolerability, as shown by the lack of thrombosis and/or embolization of the arteries.

23.4.2 Delivery to the brain: Passing the blood-brain barrier (BBB) The blood-brain barrier (BBB) is an effective obstacle for most drugs, including small molecular weight agents and macromolecules, because of the tight junctions between the endothelial lining and brain blood vessels.133 In order to achieve significant transport across the BBB, coating nanoparticles with polysorbate 80 or polysorbate with 20 polyoxyethylene units was suggested.134 An incidence of glioblastoma tumour cure in rats was observed with intravenously injected polysorbate 80 coated and doxorubicin entrapped poly(butyl 2-cyanoacrylate) nanoparticles. Polysorbates on the nanoparticles adsorb apolipoproteins B and E and are then taken up by brain capillary endothelial cells via receptor-mediated endocytosis. PEG-treated poly(alkylcyanoacylate) and PEG-hexadecylcyanoacrylate nanoparticles were shown to cross the BBB and accumulate at high densities in the brain in multiple sclerosis135 and brain tumour,136 respectively. Microparticles can also be implanted directly into the brain to bypass the BBB and problems with systemic drug administration. The safety and biocompatibility of 6 µm diameter particles composed of dipalmitoylphosphatidylcholine and chondroitin sulfate A, delivered into the cerebral parenchyma and ventricles and in intravascular injection, have been studied.137 The particles were reported to be biocompatible and caused minimal tissue injury and inflammatory reaction in in vitro and in vivo studies over 14 days. © 2008, Woodhead Publishing Limited

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Fewer studies have been reported using polymer-based particles to deliver drugs to the brain after intravenous administration.138 Apolipoprotein E was suggested to mediate the crossing of drugs through the BBB when covalently immobilized on loperamide-loaded human serum albumin nanoparticles.139 After intravenous injection in mice, the preparation was reported to induce antinociceptive effects in the tail flick test. In a similar study, apoliporotein A-I (Apo A-I) containing protamine-oligonucleotide nanoparticles was shown to be taken up and transcytosed by brain capillary endothelial cells as observed in an in vitro model of the BBB.140 Apo A-I coating also resulted in a two-fold rate of particle delivery to astrocytes of BBB in this in vitro model.

23.4.3 Targeting particles by using biological affinities Leukaemia cells expressing transferrin receptors were targeted with transferrinconjugated MS2 bacteriophage capsid proteins encapsulating antisense oligonucleotides.141 Receptor-mediated endocytosis has potential for targeting cells with antisense oligonucleotides or other nucleic acids. About eight-fold cytotoxicity was observed with encapsulated ODNs compared with control ODNs in this study. The bio recognition between lectinized drug delivery systems and glycolysated structures in the mucosal surfaces, e.g. intestine, oral cavity, respiratory track and eye, can be exploited for improved peroral therapy. The powerful mucoadhesion properties of lectin can mediate mucoadhesion, cytoadhesion, and cytoinvasion of drugs by vesicular pathways through receptor mediated endocytosis.142 Intestinal M-cells offer a portal for absorption or colloidal delivery vehicles.143 M-cell specific lectins, microbial adhesions or immunoglobilins can be used for this purpose. Various PEG-conjugated phospholipid liposomes with mean diameters of 97 nm and carrying anticancer agent pactitaxel were targeted against the folate receptors (FR) of epithelial cancer cells which over-express FR.144 FR-targeted liposomes containing the drug showed 3.8-fold greater cytotoxicity against cancer cells compared with non-targeted control liposomes. Antisense oligonucleotides (asODN) directed against c-myb or c-myc protooncogene were encapsulated in cationic liposomes and targeted against melanoma cells or neuroblastoma cells by immobilization of anti-GD2 at the particle surface.145 The asODNs delivered were reported to be effective in silencing cancer gene expression of melanoma cells. Inhibition of gene expression by asODN can occur at different levels; one best characterized involves the blocking of the translation of mRNA into protein. asODN against oncogens prevents this process, leading to cell death and slower tumour growth or regression.146 The ability of adsorbed polymers to change the biological interactions

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and distribution properties of particles with the surrounding tissue can be a valuable tool for increasing the affinity of the particle towards lymph nodes. For example, Poloxamer 407-coated nanospheres (45 nm) are controlled by a surface configuration of ethylene oxide (EO) segments which can determine the distribution of subcutaneously injected nanospheres.147 At low poloxamer surface coverage, EO tails spread laterally on a nanosphere surface and adopt a flat configuration. Such nanospheres drain rapidly from a subcutaneous injection site into initial lymphatic regions, when compared with uncoated nanospheres, which are captured by scavengers from the lymphatic nodes. When the poloxamer concentration is increased, the EO chains become more closely packed and project outward from the surface, escaping removal by the lymph node and reaching the systemic circulation. In recent studies, a more active type of drug targeting has been aimed against lymph cancers because of their higher selectivity. Bovine serum albumin (BSA) and chitosan nanospheres of mitoxantrone were studied for their therapeutic efficiency against lymph node metastases and breast cancer after local injection in rats.148 Anti-cancer loaded drugs showed a much higher drug concentration in the lymph nodes compared with the free drug. The results in an inoculated nude mice model indicated that the P388 lymph node metastases and human MCf7 breast cancer were efficiently inhibited. In a more active targeting attempt, Cirstoiu-Hapca et al. reported that antibody carrier systems show different effectiveness depending on the target.149 The effectiveness of anti-CD20 monoclonal antibody conjugated PLA nanoparticles (170 nm) against Danidi lymphoma cells (over expressing CD20 antigens) and anti-HER2 against SKOV-3 human ovarian cancer cells was compared. The selective targeting of anti-CD20-nanoparticles was detected as particles bound to and remained at the surface of lymphoma cells, while the nanoparticles with anti-HER2 were effectively internalized by ovarian cancer cells.

23.5

Delivery of bioactive agents at the target site and novel deliveries

Several techniques have been described for overcoming biological barriers, such as the blood-brain barrier, cell membranes, endosomal degradation, etc., in order to deliver bioactive agents at maximum activity. In order to see beneficial activity from the encapsulated bioactive agent, the carrier system must enter the cytoplasm and certain compartments of the cell to release the drug in active form without degradation. The entry of the particle itself, however, presents a major obstacle to drug delivery in the cell. Particles and macromolecules are internalized into the cells by a variety of mechanisms, and their intracellular fate is usually linked to the entry mechanism (Fig. 23.1).

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Particles in the size range of a few to several hundred nanometers enter the cell via uptake mechanisms including pinocytosis, non-specific endocytosis, and, for larger particles, phagocytosis.150 The exact mechanism of receptor-mediated endocytosis is not fully understood, although the receptors opsonin and lectin and scavenger receptors have been identified as part of the process.151 Endocytosis appears to be the major pathway by which oligonucleotides enter most cells, although the mechanism of entry is not fully clarified.152 Oligonucleotides are presumed to be internalized by endocytosis and somehow cross the endosomal/lysosomal membrane before being degraded.153 A current research interest in endocytosis is the stage at which the antisense oligonucleotides escape into the cytoplasm, which is crucial for genes to be delivered and express their proteins or inhibit genes (ODNs). Although the endosomal escape mechanism is not well understood, the anionic biopolymer alginate was reported to improve the transfection efficiency in cells by polyethylenimine (PEI)-alginate/DNA complexes.154 Alginate might help release the DNA complex from the endosome somehow, perhaps by rupturing of the endosome membrane. Cell penetrating peptides (CPPs) are short peptides of less than 30 amino acids that are able to penetrate cell membranes and translocate different cargoes into the cell.155 Among them, the human immunodeficiency virus (HIV) Tat-derived peptide, which is a small basic peptide, has been successfully shown to deliver small particles, including proteins, peptides and nucleic acids.156 Generally, CCPs are amphiphilic peptides, which can be internalized into cells by an internalization mechanism that requires no energy and is not necessarily receptor mediated. In order to use CPPs for drug delivery, or as part of a drug delivery vehicle, in practice, they must be small enough to synthesize, able to be coupled to different cargoes without losing their translocation properties, selective for specific cell and cell compartments, and without side effects. A few internalization studies using transport proteins have been performed in vivo in whole organisms and ex vivo in isolated tissue. In one of these studies, blood vessel endothelial tissue from mouse aorta rings was observed to take up the caveolin-1 scaffolding domain peptide attached to penetratin (protein derived from CPP), resulting in the inhibition of acetylcholineinduced vasodilation and nitric oxide production.157 In vivo use of the transduction domain of the HIV Tat protein to deliver active P-galactosidase after intraperitoneal injection was reported to increase enzyme activity in cells from all tissues.158 Recently, Sethuraman et al. showed that Tat protein conjugated to PEG, which forms the outer hydrophilic shell of PLLA core micelles, not only translocates into cytoplasm but is also seen on the surface of the nucleus, as determined by confocal microscopy.159 Several reports have indicated that

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Cellular delivery RES

Organs (1)

Endosome Lysosomal pathway Endosomal escape

Blood

(2) RES: Reticuloendothelial system (1) Receptor mediated endocytosis. (2) Non-receptor mediated endocytosis.

Biological barriers

23.1 In vivo distribution scheme of stealth protected with or without affinity ligand coupled particles and their major paths to target cell. © 2008, Woodhead Publishing Limited

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Stealth-particle with affinity ligands (▲)

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Tat-modified cationic liposomes translocated quickly and efficiently into cell cytoplasm.160,161 After internalization, Tat-modified as well as unmodified lipoplexes end up in lysosomal vesicles, indicating the involvement of clathrinmediated endocytosis. In addition, the results of cholesterol binding studies suggest that Tat-modified and unmodified cationic liposomes enter the cytoplasm mainly via cholesterol-dependent clathrin mediated endocytosis along a lipid-raft mediated pathway.161

23.6

Viral delivery systems

Viral tropism targets specific cells or tissues, and many scientific investigations have been conducted into the possibility of manipulating viruses for the treatment of diseases.162 The ability to encapsulate oligonucleotides or DNA led to the concept of using viruses as carriers for gene therapy. Retroviral vectors are one of the most powerful gene delivery vectors among viral particles. The retroviral nucleocapsid, which contains the genetic material, is wrapped in a cell-derived bilipidic membrane containing viral glycoproteins, which protects against capture by the immune system.163 In order to decrease virus-dependent cytotoxicities, continuous release systems have been proposed in place of repeated injections. The effectiveness of adenoviral vectors to deliver cytotoxic or immuno-stimulatory gene products directly to tumour sites has been studied.164 A recombinant adenoviral vector contained in PLGA microspheres was suggested as a way of decreasing inflammation because the viral vector would be released continuously in low amounts, avoiding the need for re-administration. The safety of viral vectors as drug carriers is of concern due to the possibilities of co-introduction of genetic elements from the parent virus, reexpression of viral genes, and changes in host genome structures. The protein nature of viral particles also adds to safety concerns. For that reason, it has been suggested to use parts of the virus as drug carrier instead of the whole structure. In such an approach, a chimerical gene transfer system that combines viral and nonviral features, such as using hemagglutinating virus of Japan (HVJ) fusogenic envelope combined with DNA-loaded liposomes, was found to deliver DNA, RNA, and oligonucleotides efficiently in vitro and in vivo by virus-cell fusion, which protects the contents from degradation by endosomes and lysosomes.165 A therapy which includes intratumoral injection of adenovirus particles containing the nitroreductase gene of E. coli and a systemic prodrug administration was suggested to kill cancer cells specifically in tumours.166 By this so called ‘gene directed enzyme prodrug therapy’, single viral injections into tumours followed by prodrug CB1954 treatment produced clear antitumour effects in subcutaneous xenograf models, even though virus-dependent immunological problems were detected as well. Virus-like nanoparticles (VNPs) are of interest in vaccination, gene therapy

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and drug delivery, but their potential has yet to be fully realized.167 Virus proteins that can be produced safely by recombinant DNA technology offer a new class of delivery systems with targeting advantages. The recombinantexpressed major capsid protein of polyoma virus, VP1, acts as a major ligand for certain membrane receptors to facilitate entry.168 Furthermore, the N-terminus of VP1 protein contains a DNA-binding domain and nuclear localization sequence which may provide a targeting as well as a drugbinding site for gene therapy. Similarly, MS2 bacteriophage capsid proteins were suggested to deliver ODNs as virus-like particles.141 After covalent conjugation with transferrin on the surface, the viral particles containing ODNs demonstrated increased effectiveness in killing target leukaemia cells expressing transferrin receptors. The ability to target tumours is an important goal for VNP development. Specific delivery of imaging agents or chemotherapeutics directly to tumours facilitates a greater degree of sensitivity and reduces systemic toxicity, and provides a method for monitoring tumour regression during treatment.169 Recently, Canine Parvo Virus particles (CPV-VLPs) in various human tumour cells were investigated, and it was shown that the conjugated CPV-VLPs retained their specificity for transferrin receptor-expressing cells while they did not interact with transferrin receptor-negative cells.170 Similarly, Cowpea Mosaic Virus (CPMV) conjugated with the Intron 8 region of Herstatin, which binds to Her2/Nue protein expressed in some breast cancer types, has shown that specificity of binding was preserved in vitro.171

23.7

Conclusions

Particle formulations for drug delivery provide opportunities to improve survival in humans and animals by their ability to protect bioactive reagents, permit safe delivery, modulate drug release between upper and lower indexes and reduce side effects. Importantly, the surface modification of nanoparticles presents enormous possibilities for delivering drugs to specific sites, permitting safe gene and cancer therapies. Although great advances have been made in polymer chemistry for modulating drug release rates, in stealth technology for increasing circulatory half-life, and in bio-conjugate chemistry to increase the targeting specificity of particular drug carriers, physiological and intracellular barriers remain major obstacles. In order to achieve successful targeting and delivery of bioactive agents, it is essential that the carrier system combines the full range of stealth, targeting and drug delivery properties. A more complete understanding of the molecular mechanisms behind particle translocation at the plasma membrane and transportation between cell organelles after endocytosis will be important for designing more biologically friendly particulate drug carriers based on biomimetic principles.

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Acknowledgements

E T Baran thanks the Portuguese Foundation for Science and Technology for providing him a PostDoc scholarship (SFRH/BPD/30768/2006).

23.9

References

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152 Lebedeva I, Benimetskaya L, Stein C A and Vilenchik M, ‘Cellular delivery of antisense oligonucleotides’, Eur J Pharm Biopharm, 2000, 50, 101–119. 153 Jensen K D, Kopeãková P and Kopeãek J, ‘Antisense oligonucleotides delivered to the lysosome escape and actively inhibit the hepatitis B virus’, Bioconjugate Chem, 2002, 13(5), 975–984. 154 Patnaik S, Aggarwal A, Nimesh S, Goel A, Ganguli M, Saini N, Singh Y and Gupta K C, ‘PEI-alginate nanocomposites as efficient in vitro gene transfection agents’, J Controlled Release, 2006, 114, 398–409. 155 Zorko M and Langel Ü, ‘Cell-penetrating peptides: mechanism and kinetics of cargo delivery’, Adv Drug Deliver Rev, 2005, 57, 529–545. 156 Brooks H, Lebleu B and Vives E, ‘Tat peptide-mediated cellular delivery: back to basics’, Adv Drug Deliver Rev, 2005, 57, 559–577. 157 Bucci M, Gratton J P, Rudic R D, Acevedo L, Roviezzo F, Cirino G and Sessa W C, ‘In vivo delivery of the caveolin-1 scaffolding domain inhibits nitric oxide synthesis and reduced inflammation’, Nat Med, 2000, 6, 1362–1367. 158 Schwarze S R, Ho A, Vocero-Akbani A and Dowdy S F, ‘In vivo protein transduction: delivery of a biologically active protein into the mouse’, Science, 1999, 285, 1569–1572. 159 Sethuraman V A and Bae Y H, ‘TAT peptide-based micelle system for potential active targeting of anti-cancer agents to acidic solid tumors’, J Controlled Release, 2007, 118, 216–224. 160 Hyndman L J, Lemoine L, Huang L, Porteous D J, Boyd A C and Nan X, ‘HIV-1 Tat protein transduction domain peptide facilitates gene transfer in combination with cationic liposomes’, J Controlled Release, 2004, 99, 435–444. 161 Vandenbroucke R E, De Smedt S C, Demeester J and San N N, ‘Cellular entry pathway and gene transfer capacity of TAT-modified lipoplexes’, Biochim Biophys Acta, 2007, 1768, 571–579. 162 Saini V, Zharov V P, Brazel C S, Nikles D E, Duane T and Johnson Everts M, ‘Combination of viral biology and nanotechnology: new applications in nanomedicine’, Nanomed: Nanotechnol Biol Med, 2006, 2, 200–206. 163 Lavillette D, Russell S J and Cosset F, ‘Retargeting gene delivery using surfaceengineered retroviral vector particles’, Curr Opin Biotechnol, 2001, 12, 461–466. 164 Beer S J, Hilfinger J M and Davidson B L, ‘Extended release of adenovirus from polymer microspheres: potential use in gene therapy for brain tumors’, Adv Drug Deliver Rev, 1997, 27, 59–66. 165 Kaneda Y, Nakajima T, Nishikawa T, Yamamoto S, Ikegami H N, Suzuki, Nakamura H, Morishita R and Kotani H, ‘Hemagglutinating Virus of Japan (HVJ) envelope vector as a versatile gene delivery system’, Mol Ther, 2002, 6(2), 219–226. 166 Djeha A H, Thomson T A, Leung H, Searle P F, Young L S, Kerr D J, Harris P A, Mountain A and Wrighton C J, ‘Combined adenovirus-mediated nitroreductase gene delivery and CB1954 treatment: a well-tolerated therapy for established solid tumors’, Mol Ther, 2001, 3(2), 233–240. 167 Pattenden L K A, Middelberg P J, Niebert M and Lipin D I, ‘Towards the preparative and large-scale precision manufacture of virus-like particles’, Trends in Biotechnol, 2005, 23(10), 523–529. 168 Georgens C, Weyermann J and Zimmer A, ‘Recombinant virus like particles as drug delivery system’, Curr Pharm Biotechnol, 2005, 6(1), 49–55. 169 Manchester M and Singh P, ‘Virus-based nanoparticles (VNPs): Platform technologies for diagnostic imaging’, Adv Drug Deliver Rev, 2006, 58, 1505–1522.

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170 Singh P, Destito G, Schneemann A and Manchester M, ‘Canine parvovirus-like particles, a novel nanomaterial for tumor targeting’, J Nanobiotechnol, 2006, 4, 2– 12. 171 Chatterji A, Ochoa W, Shamieh L, Salakian P, Wong S M, Clinton G, Ghosh P, Lin T W and Johnson J E, ‘Chemical conjugation of heterologous proteins on the surface of cowpea mosaic virus’, Bioconjugate Chem, 2004, 15, 807–813.

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24 Thiolated chitosans in non-invasive drug delivery A. B E R N K O P - S C H N Ü R C H, Leopold-Franzens University, Austria

24.1

Introduction

Chitosan is a polysaccharide consisting of copolymers of glucosamine and N-acetylglucosamine. It is obtained by alkaline deacetylation of chitin, which is the second most abundant polysaccharide in nature. Shell wastes of crab, shrimp and lobster are the main industrial sources of chitin.1 The primary amino group of chitosan accounts for the possibility of relatively easy chemical modification and salt formation with acids. At acidic pH the amino groups are protonated, which promotes solubility, whereas chitosan is insoluble at alkaline and neutral pH.2 Because of its favorable properties, such as biocompatibility, enzymatic biodegradability and non-toxicity,1 chitosan has received considerable attention as an excipient in drug delivery systems. Since 2002 chitosan has a monography in the European Pharmacopoeia. Because of its mucoadhesive and/or permeation enhancing properties it is utilized in various fields of pharmaceutical technology including the formulation of controlled release dosage forms, such as tablets, gels, micro- and nanoparticles for oral, nasal, ocular and buccal drug delivery.3–6 In addition, chitosan shows promising features in non-viral gene delivery.7,8 The mucoadhesive and permeation enhancing properties can be further improved by the immobilization of thiol groups on the polymeric backbone of chitosan. To date, various different thiolated chitosan derivatives have been synthesized including chitosan-thioglycolic acid conjugates,9–11 chitosancysteine conjugates,12 chitosan-N-acetylcysteine conjugates,13 chitosan-thioethyl-amidine conjugates (chitosan-TEA),14 chitosan-4-thio-butyl-amidine (chitosan-TBA)15,16 and chitosan-glutathione conjugates.17 Apart from improved mucoadhesive and permeation enhancing properties, thiolated chitosans show also strong cohesive properties, in situ gelling11 and efflux pump inhibitory properties.18,19 It is the aim of this chapter to provide an overview about different thiolated chitosans that have been generated so far, and their characterization and 624 © 2008, Woodhead Publishing Limited

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optimization utilizing various in vitro test systems. The performance of thiolated chitosans in vivo and potential future applications in drug delivery will be discussed.

24.2

Thiolated chitosans

The primary amino group at the 2-position of the glucosamine subunits of chitosan is the main target for the immobilization of thiol groups. As outlined in Fig. 24.1 sulfhydryl bearing agents can be covalently attached to this primary amino group via the formation of amide or amidine bonds. In the case of the formation of amide bonds, the carboxylic acid group of the ligands cysteine, N-acetylcysteine, glutathione and thioglycolic acid react with the primary amino group of chitosan mediated by a water soluble carbodiimide. The formation of disulfide bonds by air oxidation during the synthesis is avoided by performing the process at pH 5. At this pH-range the concentration of thiolate-anions, representing the reactive form for oxidation of thiol groups, is marginal, and the formation of disulfide bonds can be almost excluded. Alternatively, the coupling reaction can be performed under inert conditions. In the case of the formation of amidine bonds, 2-iminothiolane can be used as a coupling reagent. It offers the advantage of a simple one step coupling reaction. In addition, the thiol group of the reagent is protected towards oxidation because of the chemical structure of the reagent. Although the reaction is very simple leading to a comparatively high degree of derivatization, the resulting conjugate has the drawback of chemical instability. This chemical instability of chitosan-TBA led to the development of chitosanTEA.14 The formation of cyclic non-thiol products during synthesis and storage is excluded in contrast to chitosan-TBA conjugate where 2iminothiolane is used as a modifying reagent.15 The amount of immobilized thiol groups in reduced and oxidized form can be determined via Ellman’s reagent11 with and without previous quantitative reduction of disulfide bonds with borohydride.20

24.3

Properties of thiolated chitosans

24.3.1 Mucoadhesive properties The improved mucoadhesive properties of thiolated chitosans are explained by the formation of covalent bonds between thiol groups of the polymer and cysteine - rich subdomains of glycoproteins in the mucus layer. Ultimate evidence for thiol/disulfide exchange reactions has been provided by Leitner et al.21 These covalent bonds are much stronger than non-covalent bonds, such as ionic interactions of chitosan with anionic substructures of the mucus

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SH HS H2N H

H

NH

NH

H H

OH H

H O

O

H

CH2OH

H

Chitosan-Thioglycolic Acid

NH2 Cl– H

H O

H O

CH2OH Chitosan-N-Acetyl-Cysteine

24.1 Substructure of thiolated chitosans. © 2008, Woodhead Publishing Limited

Glutathione

NH

H

NH

H OH H

O

O

SH

O NH

H

Chitosan-4-Thio-Butyl-Amidine

HS

H H

H CH2OH

H OH

O

O

HS Ac–HN

NH H

OH H

CH2OH

Chitosan-Cysteine

H

H H

H OH

NH2 Cl–

O

H OH

H O

H O

CH2OH Chitosan-Thio-Ethyl-Amidine

H

H O

O

CH2OH Chitosan-Glutathione

Natural-based polymers for biomedical applications

SH O

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Table 24.1 Rank order of the ten most mucoadhesive polymers determined via the rotating cylinder method (adapted from Ref. 23) Polymer

pH

Chitosan-TBA Chitosan-TBA Polyacarbophil–cysteine Chitosan-TBA Polyacrylicacid (450 kDA)-cysteine Hydroxypropyl-cellulose Carbopol 980 Carbopol 974 Polycarbophil Carbopol 980

pH pH pH pH pH pH pH pH pH pH

Time in hours; means ± SD (n = 3–5) 3 lyophilized 6.5 precipitated 3 lyophilized 6.5 lyophilized 3 lyophilized 7 precipitated 7 precipitated 7 precipitated 7 precipitated 3 lyophilized

161.2 ± 7.2 40.4 ± 2.1 26.0 ± 0.9 20.4 ± 1.5 19.4 ± 0.8 15.2 ± 0.4 12.5 ± 0.9 10.3 ± 0.9 10.2 ± 0.8 9.8 ± 0.2

layer. This theory was supported by the results of tensile studies with tablets of thiolated chitosans, which demonstrated a positive correlation between the degree of modification with thiol bearing moieties and the adhesive properties of the polymer.10,16,22 These findings were confirmed by another in vitro mucoadhesion test system, where the time of adhesion of tablets on intestinal mucosa is determined. The contact time of the thiolated chitosan derivatives increased with increasing amounts of immobilized thiol groups.10,15 With chitosan-thioglycolic acid conjugates a 5 to 10-fold increase in mucoadhesion in comparison to unmodified chitosan was achieved. The mucoadhesive properties of chitosan-TBA conjugates were even further improved. One explanation for this phenomenon can be given by the theory that chitosan-TBA conjugates have additionally increased mucoadhesive properties due to improved ionic interactions between the additional cationic amidine substructure of the conjugate as illustrated in Fig. 24.1 and anionic substructures within the mucus layer. Tensile studies with chitosan-TBA conjugates of low, medium and high molecular mass (150, 400 and 600 kDa) furthermore indicated that medium molecular mass thiolated chitosans display the relatively highest mucoadhesiveness. Utilizing a medium molecular mass chitosan-TBA conjugate displaying 264 µM thiol groups per g polymer consequently led to a more than 100-fold improvement in mucoadhesion in comparison with unmodified chitosan. This represents the greatest progress made so far in the development of mucoadhesive polymers.16 Recently, Grabovac et al. determined the mucoadhesive properties of all polymers which are referred in the literature to be mucoadhesive, resulting in a rank order from 1–67. In Table 24.1 the ten most mucoadhesive polymers are ranked. Another likely mechanism being responsible for the improved mucoadhesive properties of thiomers is based on their in situ cross-linking properties. During and after the interpenetration process, which were verified for mucoadhesive polymers such as poly(acrylic acid) recently,24

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disulfide bonds are formed within the thiomer itself leading to additional anchors via chaining up with the mucus gel layer. It is similar to the mechanism on which the adhesive properties of most adhesives are based. That is a penetration of the adhesive into a certain surface structure followed by a stabilization process of the adhesive. In the case of superglues, for instance, monomeric cyanoacrylates penetrate into raw surfaces followed by a polymerization process. Thiolated polymers display in situ gelling properties due to the oxidation of thiol groups at physiological pH-values, which results in the formation of inter- and intramolecular disulfide bonds.

24.3.2 In-situ gelling properties Rapid clearance from the site of drug action is one important factor that limits the efficacy of drugs administered to the nasal, ocular and vaginal mucosa. It is widely accepted that limiting the clearance by increasing the viscosity of a drug formulation will result in an increased bioavailability of these drugs. The formation of a gel at the site of drug delivery combines the advantages of a solution, which can be easily administered, with the favorable viscoelastic properties of a gel providing a prolonged residence time of the formulation. The sol–gel transition occurs in the physiological environment as a result of physicochemical changes such as changes in temperature,25,26 in the pH27 or in electrolyte concentration.28,29 Thiolated chitosans display in situ gelling properties due to disulfide bond formation at physiological pHvalues. This cross-linking process can be observed at pH levels above 5.11 The in situ gelling behavior of thiolated chitosans was characterized in vitro by rheological measurements. The sol–gel transition of thiolated chitosans at pH 5.5 was completed within two hours, when highly crosslinked gels have been formed. In parallel, a significant decrease in the thiol group content of the polymers can be observed, indicating the formation of disulfide bonds. 11,15 Rheological investigation of thiolated chitosans furthermore demonstrated a clear correlation between the total amount of polymer-linked thiol groups and the increase in elasticity of the formed gel. The more thiol groups are immobilized on chitosan, the higher is the increase in elastic modulus G′, which is illustrated in Fig. 24.2.11,15 According to this, thiolated chitosans seem to be promising new excipients for liquid or semisolid formulations stabilizing themselves once having been applied on the site of drug delivery.

24.3.3 Permeation enhancing properties In 1994 the permeation enhancing capabilities of chitosan were shown for the first time.4 Chitosan is able to enhance the paracellular route of uptake, which is the favorable route for hydrophilic compounds such as therapeutic

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1,E+03

G′ (Pa)

1,E+02

1,E+01

1,E+00

1,E–01 0

1

2 4 Time [hour]

6

24.2 Increase in G′ of a 1.5% (m/v) chitosan-TBA conjugate gel at pH 5.5 and 37°C as a function of time. Adapted from Bernkop-Schnürch et al.15

peptides and proteins.30–32 The mechanism underlying this permeation enhancing effect seems to be based on the positive charges of the polymer, which interact with the cell membrane resulting in a structural reorganization of tight junction-associated proteins.33 In the presence of the mucus layer, however, this permeation enhancing effect is comparatively lower, as chitosan cannot reach the cell surface because of size limited diffusion and/or competitive charge interactions with mucins.34 Nevertheless, these results obtained on Caco-2 cell monolayers could be confirmed by in vivo studies, showing an enhanced intestinal absorption of the peptide drug buserelin in rats due to the co-administration of chitosan hydrochloride.35 This permeation enhancing effect of chitosan can be strongly improved by the immobilization of thiol groups on the polymer. The uptake of fluorescence labeled bacitracin, for instance, was 1.6-fold improved utilizing 0.5% of chitosan-cysteine conjugate instead of unmodified chitosan.12 In another study Langoth et al. could enhance the permeation of a therapeutic peptide in the presence of thiolated chitosan even more than 40-fold in comparison to unmodified chitosan.36 Results of this study are shown in Fig. 24.3. Such in vitro results could meanwhile be verified in numerous in vivo studies demonstrating the potential of thiomers in absorption enhancement of orally given drugs. Studies in pigs, for instance, demonstrated a significantly improved uptake of various model peptide drugs when incorporated in a thiolated chitosan.37,38

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PACAP [µg/ml]

30 25 20 15 10 5 0 0

30

60

90

120 150 Time [min]

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24.3 Permeation enhancing effect of 1% (m/v) chitosan-TBA conjugate with 2% (m/v) glutathione (X), 1% (m/v) chitosan-TBA conjugate (❍) and 1% (m/v) unmodified chitosan (䊐 ) on the uptake of a peptide drug from the porcine buccal mucosa. Adapted from Langoth et al.36

The mechanism being responsible for this improved permeation enhancement has been ascribed to be likely based on the inhibition of protein tyrosine phosphatase. This enzyme seems to be involved in the opening and closing process of the tight junctions. Protein tyrosine phosphatase is responsible for the dephosphorylation of tyrosine subunits of occludin, representing a major transmembrane protein in the region of the tight junctions. When tyrosine subunits of occludin are dephosphorylated, the tight junctions are closed. In contrast, when these tyrosine subunits are phosphorylated, the tight junctions are opened. The inhibition of protein tyrosine phosphatase by compounds such as pervanadate, phenylarsine oxide or reduced glutathione leads consequently to phosphorylation and opening of tight junctions.39–41 In contrast to the stable protein tyrosine phosphatase inhibitors pervanadate and phenylarsine oxide, the inhibitory effect of glutathione is quite poor as it is rapidly oxidized on the cell surface losing its inhibitory activity.42 Due to the combination of reduced gutathione with thiolated chitosans, however, this oxidation of the inhibitor on the membrane can be avoided, as thiomers are capable of reducing oxidized glutathione.39

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24.3.4 Efflux pump inhibitory properties Recently Werle and Hoffer revealed efflux pump inhibitory capability of thiolated chitosans.18 The transmucosal transport of the P-gp substrate rhodamine 123 was strongly improved in the presence of thiolated chitosan. These in vitro results could meanwhile be confirmed by in vivo studies in rats. Föger et al. showed that the oral bioavailability of rhodamine 123 is even 3.0-fold improved when this model P-gp substrate is embedded in thiolated chitosan minitablets given orally to rats. Results of this study are shown in Fig. 24.4.19 The postulated mechanism of efflux pump inhibition is based on an interaction of thiomers with the channel forming transmembrane region of P-gp. P-gp exhibits 12 transmembrane regions forming a channel through which substrates are transported outside of the cell. The transmembrane regions 2 and 11 exhibit on position 137 and 956, respectively, a cysteine subunit.43 Thiomers seem to enter in the channel of P-gp forming subsequently one or two disulfide bonds with one or both cysteine subunits located within the channel. The theory is supported by the size dependent activity of thiomers and the observation that corresponding unthiolated polymers show no or significantly lower efflux-pump inhibitory properties. The theory is additionally supported by the observation that the efflux pump inhibitory effect of thiomers is only to a minor extent influenced by the type of polymer backbone. For instance, thiolated chitosans show a similar effect as obtained by using thiolated polyacrylates.44 Apart from their oral use, thiomers might also be helpful in order to improve the efficacy of parenterally administered anticancer drugs.

Rhodamine 123 [ng/mL]

50

40

30

20

10

0 0

2

4

6 Time [hour]

8

10

12

24.4 Pharmacokinetic of orally administered rhodamine 123 (Rho-123) in rats; rhodamine 123 given in solution (X); rhodamine 123 given in the presence of thiolated chitosan (♦ ♦ ). Adapted from Föger et al.19

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24.3.5 Transfection enhancing properties Site-specific gene delivery requires vectors that are stable in the extracellular environment, enabling systemic delivery, but after arrival in the target cell the vector should release the contained nucleic acid in a suitable form for transcription or translation. One biological trigger mechanism is the different reducing environment in the plasma (estimated 1–2 µM glutathione) and within the cell (estimated 1–20 mM glutathione). Because of this reducing milieu within the cell, disulfide bonds as linkers between the hydrophilic polymer coating and the DNA complex offer the possibility to be hydrolyzed in the endosome or cytoplasma, leading to a release of the DNA complex from the vector, as illustrated in Figure 24.5. Vectors with different sulfhydryl bearing agents for intraparticular disulfide binding showed significantly higher transfection rates in vitro than control vectors with thioether bindings.45 Another study could show that thiolated gelatin nanoparticles were significantly more effective in transfecting NIH-3T3 cells than other unthiolated carriers examined.46 Because of this advantage, our research group decided to experiment with sulfhydryl bearing chitosan itself. Martien et al. could show that when thioglycolic acid (TGA) was immobilized to chitosan/pDNA nanoparticles, transfection rates were even 5-fold higher in Caco-2 cells than those gained with unmodified chitosan/DNA nanoparticles. Especially when

Extracellular (oxidizing conditions)

Cytoplasma (reducing conditions)

SH S S

HS

SH

HS S S

SH

SH

SH

S S SH SS

SS

HS SH

SH SH

24.5 Because of the reducing milieu within the cell, disulfide bonds as linkers between the hydrophilic polymer coating and the DNA complex offer the possibility to be hydrolyzed in the endosome or cytoplasma, leading to a release of DNA (dotted circles) from the polymeric carrier systems.

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chitosan-TGA/pDNA nanoparticles were crosslinked by oxidation with H2O2, expression of green fluorescence protein could be enhanced compared to non-crosslinked nanoparticles.47 Chitosan-TBA nanoparticles were found to increase transfection rates in Caco-2 cells as well. The relative highest efficiency for chitosan-TBA/pDNA nanoparticles was reached in combination with glycerol shock solution.48 In another study N-acetylcysteine-cysteine was covalently attached to chitosan. Nanoparticles comprising this thiomer and pDNA showed significantly improved transfection and expression rates of transgenes as well.49

24.4

Drug delivery systems

24.4.1 Gels The great advantage of thiolated chitosans for gel formulations is their mucoadhesive and in situ gelling properties.15 Strong mucoadhesive properties are useless if the adhesive bond fails within the gel formulation itself rather than between the gel and the mucus gel layer covering mucosal membranes. Due to the in situ crosslinking properties of thiolated chitosans, however, this shortcoming can be overcome.

24.4.2 Matrix tablets Thiolated chitosan matrix tablets are useful for intraoral, peroral, ocular and vaginal – local or systemic drug delivery. Due to the in situ cross-linking properties of thiolated chitosans the cohesiveness and subsequently the stability of the swollen carrier matrix can be guaranteed.50 When attached in dry form the mucoadhesion of matrix tablets is additionally improved by an augmented interpenetration process depending on the swelling behavior of the delivery system. In order to make use of this simple adhesion by a hydration process in the case of oral delivery, the matrix tablets can be enteric coated. If an adhesion is achieved already in the stomach, coating with a triglyceride is sufficient in order to avoid an unintended adhesion in the oral cavity or oesophagus.51 Moreover, matrix tablets comprising a thiolated chitosan offer the advantage that a controlled drug release can be achieved out of this type of dosage form, which can already be demonstrated for numerous drugs [e.g. 36,52,53] Langoth et al., for instance, could show that by simply homogenizing the thiomer with the drug of choice and compressing tablets out of it leads to delivery systems which can guarantee even a zero order release profile for several hours.36 The drug release rate is thereby mainly controlled by a hydration and diffusion process. In addition drug release can be controlled by making use of ionic interactions between the cationic chitosan and anionic drugs.

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24.4.3 Micro- and nanoparticles Because of their relative small size micro- and nanoparticles show a prolonged gastrointestinal residence time even without any mucoadhesive properties by diffusing into the mucus gel layer.54 Coupe et al., for instance, could demonstrate that particulate delivery systems display a more prolonged gastrointestinal transit time compared to single-unit dosage forms.55 In order to further improve the residence time of drug delivery systems on mucosal membranes, both approaches: mucoadhesive polymers (I) and micro-/ nanoparticles (II) were consequently combined. Micro- and nanoparticles based on unmodified chitosans, however, disintegrate very rapidly, unless multivalent anions such as sulfate or tripolyphosphate (TPP) are added leading to stabilization via an ionic cross-linking process.56 Due to the addition of such ionic cross-linkers, on the other hand, the mucoadhesive properties are strongly reduced. Conversely, due to the immobilization of thiol groups on modified polymers, their mucoadhesive properties are even further improved, although micro- and nanoparticles being based on thiolated polymers do not disintegrate. Because of the formation of disulfide bonds within the polymeric network, the particles are stabilized.57 In general the preparation of thiolated chitosan nanoparticles is based on the following steps. In the first step thiolated chitosan is ionically gelated with tripolyphosphate or sulfate in aqueous solution forming submicron particles and microparticles, respectively. In the next step thiol groups in and on the particles are partially oxidized forming stabilizing inter- and intramolecular disulfide bonds. As the degree of oxidation can be controlled during the production process, the share of thiol and disulfide groups can be adjusted on demand. Thereafter the polyanions are removed. Utilizing this novel preparation method as illustrated in Fig. 24.6, stable particles of a mean size in the range from 100 nm up to 10 µm can be produced. Consequently, also a controlled drug release out of the thiomer micro- and nanoparticles can be provided. Neither ionically nor covalently crosslinked particles were degraded by lysozyme under physiological conditions.57 Results obtained with such particles demonstrated significantly improved mucoadhesive properties. The more thiol groups were oxidized within the particles, however, the lower was the improvement in mucoadhesive properties. Nevertheless, even when 91% of all thiol groups on the nanoparticles were oxidized, their mucoadhesive properties were still twice as high as that of unmodified nanoparticles.58

24.5

In vivo performance

The potential of thiolated chitosans for the oral administration of hydrophilic macromolecules could meanwhile be shown by various in vivo

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PA2–

+

NH3

+

NH3

+

+

NH3

NH3

635

PA2–

+

NH3

+

NH3

PA2–

SH SH

+

SH

NH3

SH

Ox

SH

S –PA2–

S

S

S

S

PA2–

PA2–

SH

S S S

S

S

PA

2–

24.6 Illustration of thiolated chitosan nanoparticle preparation. Due to the addition of di- or trivalent polyanions (PA) to thiolated chitosan solution nanoparticles are formed. In the following oxidation step these nanoparticles are stabilized via inter- and intrachain disulfide bond formation. Finally the polyanions are removed via dialysis.

studies.38,51,59–60 As model drug, for instance, salmon calcitonin was utilized, which is a peptide drug of cationic net charge and a molecular mass of 3.2 kDa. Salmon calcitonin is used for the treatment of chronic bone diseases. It is currently marketed in nasal spray and injectable forms, both having the drawback of a low patient acceptance. A higher patient compliance should be achieved by the application of an oral delivery system for this drug. Until now the oral bioavailability reached, however, is too low to permit therapeutic employment. For this reason the peptide was regarded as a challenging model drug for testing the potential of thiolated chitosans. A drug delivery system comprising thiolated chitosan and salmon calcitonin was generated. Compounds were simply homogenized and directly compressed into tablets. Tablets were orally given to rats and the plasma calcium level was monitored as a function of time. Results showed no significant reduction of the plasma calcium level caused by salmon calcitonin, which was orally given in solution. Furthermore, no significant effect was observed after oral administration of tablets comprising the therapeutic peptide and unmodified chitosan, although the native polymer is reported to be mucoadhesive and to exhibit a permeation enhancing effect for hydrophilic macromolecules. In contrast, as shown in Fig. 24.7, in the presence of thiolated chitosan a decrease in plasma calcium level of more than 5% for several hours was observed.51

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Plasma calcium level in %

100

95

90

85 0

4

8

12 Time (h)

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24.7 Decrease in plasma calcium level as a biological response for the salmon calcitonin bioavailability in rats after oral administration of thiolated chitosan/glutathione tablets targeted to the small intestine (–); of unmodified chitosan tablets targeted to the small intestine (䊊); of thiolated chitosan/glutathione tablets targeted to the stomach (*); and of unmodified chitosan tablets targeted to the stomach (◆). Adapted from Guggi et al.51

In another study a sustained buccal delivery system for pituitary adenylate cyclase-activating polypeptide (PACAP) representing a promising therapeutic peptide for the treatment of type 2 diabetes has been generated. Chitosan-4thiobutylamidine conjugate was homogenized with the enzyme inhibitor and permeation mediator glutathione, Brij 35 and PACAP.36 The mixture was lyophilized and compressed into flat-faced disks and additionally coated on one side with palm wax. Bioavailability studies were performed in pigs by buccal administration of this test formulation showing an absolute bioavailability of 1%, whereas PACAP did not reach the systemic circulation when administered via tablets consisting of unmodified chitosan and PACAP or of unmodified chitosan, Brij 35, and PACAP. Results of this study are shown in Fig. 24.8.37 Krauland et al. developed a microparticulate delivery system based on a thiolated chitosan conjugate for the nasal application of insulin. A mixture of the chitosan-TBA conjugate, insulin and the permeation mediator reduced glutathione were formulated to microparticles via the emulsification solvent evaporation technique. Nasal administered chitosan-TBA-insulin microparticles led to an absolute bioavailability of 7.24+/–0.76% in conscious rats. In contrast,

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250

PACAP [pg/ml]

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150

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50

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1

2

3 4 Time [h]

5

6

7

24.8 Plasma concentration of pituitary adenylate cyclase-activating polypeptide (PACAP) after buccal administration of PACAP/chitosanTBA tablets to pigs. Tablets were removed after six hours. Indicated values are means of four experiments. Adapted from Langoth et al.37

chitosan-insulin microparticles and mannitol-insulin microparticles exhibited an absolute bioavailability of 2.04+/–1.33% and 1.04+/–0.27%, respectively.61 Because of these results, microparticles comprising chitosan-TBA and reduced glutathione seem to represent a useful formulation for the nasal administration of peptides. Focusing on the efflux pump inhibitory properties of thiolated chitosans a direct in vivo comparison of delivery systems based on Pluronic P85, Myrj 52 and chitosan-4-thio-butyl-amidine using rhodamine-123 (Rho-123) as representative P-gp substrate has been performed in rats recently. Entericcoated tablets based on Pluronic P85, Myrj 52 or chitosan TBA/GSH increased the area under the plasma concentration time curve of Rho-123 1.6-fold, 2.4fold, 4.3-fold, respectively, in comparison to control only. This in vivo study, being illustrated in Fig. 24.9, showed that polymeric P-gp inhibitors and especially the delivery system based on thiolated chitosan significantly increased the oral bioavailability of P-gp substrate Rho-123.62 Results of this study were confirmed by another in vivo study showing in rats that Chitosan TBA/GSH tablets increase the area under the plasma concentration time curve of Rho-123 by 217% in comparison to buffer control and by 58% in comparison to unmodified chitosan.19

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10

12

24.9 Pharmacokinetic of rhodamine 123 (Rho-123) having been orally co-administrated with Pluronic P85 (∆), Myrj 52 (䊐) and chitosan-4thio-butyl-amidine (䊊) to rats; adapted from Föger.62

Studies performed in collaboration of Serono and the University of Vienna performed in pigs demonstrated a relative bioavailability of 34.4% for the subcutaneous administration of the peptide drug antide. Oral administration of antide in solution led to no detectable concentrations of the drug in plasma at all. In contrast, administering antide incorporated in thiolated chitosan resulted in a significant uptake of the peptide. The absolute and relative bioavailability was determined to be 1.1% and 3.2%, respectively. Results are shown in Fig. 24.10.38

24.6

Conclusion

Due to the covalent attachment of various reagents bearing sulfhydryl functions to chitosan various properties of this polymer can be significantly improved. Thiolated chitosans show comparatively much stronger mucoadhesive, cohesive and in situ gelling properties. Moreover a comparatively more pronounced permeation enhancing effect is provided, which can be further raised by the combination of thiolated chitosans with the permeation mediator glutathione. In addition, thiolated chitosans display efflux pump inhibitory properties. Due to these advantageous features thiolated chitosans have been successfully used for the non-invasive administration of many challenging drugs. Moreover,

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30

Antide [ng/ ml]

25 20 15 10 5 0 0

6

12 Time [hour]

18

24

24.10 Concentration profile of orally administered antide (2 mg/kg) in plasma of two pigs. The drug has been embedded in thiolated chitosan tablets. The administration of antide (2 mg/kg) in solution did not lead to an uptake of the peptide drug at all. × = Pig 1, ∆ = Pig 2. Adapted from Bernkop-Schnürch et al. [38].

thiolated chitosans can be used for gene delivery. They seem to represent a promising new generation of polymeric excipients.

24.7

References

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Natural-based polymers for biomedical applications and in vitro evaluation of chitosan-thioglycolic acid conjugates, Biomaterials, 22(2001), 2345–2352. Hornof M D, Kast C E and Bernkop-Schnürch A, In vitro evaluation of the viscoelastic behavior of chitosan – thioglycolic acid conjugates, Eur J Pharm Biopharm, 55(2003), 185–190. Bernkop-Schnürch A, Brandt U M and Clausen A E, Synthesis and in vitro evaluation of chitosan-cysteine conjugates, Sci Pharm, 67(1999), 196–208. Pichayakorn W, Ritthidej G C, Loretz B, Martien R, Schmitz Th, Kafedjiiski K and Bernkop-Schnürch A, Oral Gene Delivery: Comparison of Different Thiolated Chitosans as pDNA Carrier Matrix, Proceed, 33rd Ann, Meeting and Exp. of the Controlled Release Society 2006 Vienna. Kafedjiiski K, Hoffer M, Werle M and Bernkop-Schnürch A, Improved synthesis and in vitro characterization of chitosan-thioethylamidine conjugate, Biomaterials, 27(2006), 127–135. Bernkop-Schnürch A, Hornof M and Zoidl T, Thiolated polymers – thiomers: modification of chitosan with 2-iminothiolane, Int J Pharm, 260(2003), 229–237. Roldo M, Hornof M, Caliceti P and Bernkop-Schnürch A, Mucoadhesive thiolated chitosans as platforms for oral controlled drug delivery: Synthesis and in vitro evaluation, Eur J Pharm Biopharm, 57(2004), 115–121. Kafedjiiski K, Föger F, Werle M and Bernkop-Schnürch A, Synthesis and in vitro evaluation of a novel chitosan-glutathione conjugate, Pharm. Res., 22(2005), 1480– 1488. Werle M and Hoffer M, Glutathione and thiolated chitosan inhibit multidrug resistance P-glycoprotein activity in excised small intestine, J Control Release, 111(2006), 41– 46. Föger F, Schmitz Th and Bernkop-Schnürch A, In vivo evaluation of an oral delivery system for P-gp substrates based on thiolated chitosan, Biomaterials, 27(2006), 4250–4255. Habeeb A F, A sensitive method for localization of disulfide containing peptides in column effluents, Anal Biochem, 56(1973), 60–65. Leitner V M, Walker G F and Bernkop-Schnürch A, Thiolated polymers: Evidence for the formation of disulphide bonds with mucus glycoproteins, Eur J Pharm Biopharm, 56(2003), 207–214. Albrecht K, Zirm E J, Palmberger T F, Schlocker W and Bernkop-Schnürch A, Preparation of thiomer microparticles and in vitro evaluation of parameters influencing their mucoadhesive properties, Drug Dev Ind Pharm, 32(2006), 1149–1157. Grabovac V, Guggi D and Bernkop-Schnürch A, Comparison of the mucoadhesive properties of various polymers, Adv Drug Del Rev, 57(2005), 1713–1723. Imam M E, Hornof M, Valenta C and Bernkop-Schnürch A, Evidence for the interpenetration of mucoadhesive polymers into the mucus gel layer, STP Pharma, 13(2003), 171–176. Edsman K, Carlfors J and Petersson R, Rheological evaluation of poloxamer as an situ gel for ophthalmic use, Eur J Pharm Sci, 6(1998), 105–112. Bromberg L E, Enhanced nasal retention of hydrophobically modified polyelectrolytes, J Pharm Pharmacol, 53(2001), 109–114. Gurny R, Ibrahim H and Buri P, The development and use of in situ formed gels, triggered by pH, in: Edman P (Ed.) Biopharm Ocul Drug Delivery, CRC Press, Inc., Boca Raton, FL, 1993, pp. 81–90.

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28 Deasy P B and Quigley K J, Rheological evaluation of deacetylated gellan gum (Gelrite) for pharmaceutical use, Int J Pharm, 73(1991), 117–123. 29 Paulsson M, Hägerström H and Edsman K, Rheological studies of the gelation of deacetylated gellan gum (Gelrite) in physiological conditions, Eur J Pharm Sci, 9(1999), 99–105. 30 Artursson P, Lindmark T, Davis S S and Illum L, Effect of chitosan on the permeability of monolayers of intestinal epithelial cells (Caco-2), Pharm Res, 11(1994), 1358– 1361. 31 Borchard G, Luessen H L, de Boer A G, Verhoef J C, Lehr C-M and Junginger H E, The potential of mucoadhesive polymers in enhancing intestinal peptide drug absorption. III: Effects of chitosan-glutamate and carbomer on epithelial tight junctions in vitro, J Control Rel, 39(1996), 131–138. 32 Dodane V, Amin Khan M and Merwin J R, Effect of chitosan on epithelial permeability and structure, Int J Pharm, 182(1999), 21–32. 33 Schipper N G M, Olsson S, Hoogstraate J A, deBoer A G, Varum K M and Artursson P, Chitosans as absorption enhancers for poorly absorbable drugs, 2: Mechanism of absorption enhancement, Pharm Res, 14(1997) 923–929. 34 Schipper N G M, Varum K M, Stenberg P, Ocklind G, Lennernäs H and Artursson P, Chitosans as absorption enhancers for poorly absorbable drugs. 3: Influence of mucus on absorption enhancement, Eur J Pharm Sci, 8(1999), 335–343. 35 Luessen H L, de Leeuw B J, Langemeyer M W, de Boer A G, Verhoef J C and Junginger H E, Mucoadhesive polymers in peroral peptide drug delivery. VI. Carbomer and chitosan improve the intestinal absorption of the peptide drug buserelin in vivo, Pharm Res, 13(1996), 1668–1672. 36 Langoth N, Kalbe J and Bernkop-Schnürch A, Development of a mucoadhesive and permeation enhancing buccal delivery system for PACAP (pituitary adenylate cyclaseactivating polypeptide), Int J Pharm, 296(2005), 103–111. 37 Langoth N, Kahlbacher H, Schöffmann G, Schmerold I, Schuh M, Franz S, Kurka P and Bernkop-Schnürch A, Thiolated chitosans: design and in vivo evaluation of a mucoadhesive buccal peptide drug delivery system, Pharm Res, 23(2006), 573–579. 38 Bernkop-Schnürch A, Pinter Y, Guggi D, Kahlbacher H, Schoffmann G, Schuh M, Schmerold I, Del Curto M D, D’Antonio M, Esposito P and Huck C, The use of thiolated polymers as carrier matrix in oral peptide delivery – proof of concept, J Control Release, 106(1–2)(2005), 26–33. 39 Clausen A E, Kast C E and Bernkop-Schnürch A, The role of glutathione in the permeation enhancing effect of thiolated polymers, Pharm Res, 19(2002), 602– 608. 40 Barrett W C, DeGnore J P, Konig S, Fales H M, Keng Y F, Zhang Z Y, Yim M B and Chock P B, Regulation of PTP1B via glutathionylation of the active site cysteine 215, Biochemistry, 38(1999) 6699–6705. 41 Staddon J M, Herrenknecht K, Smales C and Rubin L L, Evidence that tyrosine phosphorylation may increase tight junction permeability, J Cell Sci, 108(1995), 609–619. 42 Grafstrom R, Stead A H and Orrenius S, Metabolism of extracellular glutathione in rat small-intestinal mucosa, Eur J Biochem, 106(1980), 571–577. 43 Ferté J, Analysis of the tangled relationships between P-glycoprotein-mediated multidrug resistance and the lipid phase of the cell membrane, Eur J Biochem, 267, (2000), 277–294. 44 Greindl M, Föger F and Bernkop-Schnürch A, In vivo evaluation of thiolated

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Natural-based polymers for biomedical applications poly(acrylic acid) as a drug absorption modulator for MRP2 efflux protein pump substrates, Innano Congress Innsbruck Austria (2006) p. 21. Carlisle R C, Etrych T, Briggs S S, Preece J A, Ulbrich K and Seymour L W, Polymer-coated polyethylenimine/DNA complexes designed for triggered activation by intracellular reduction, J Gene Med, 6(3)(2004), 337–344. Kommareddy S and Amiji M, Preparation and evaluation of thiol-modified gelatin nanoparticles for intracellular DNA delivery in response to glutathione, Bioconjug Chem, 16(6)(2005), 1423–1432. Martien R, Loretz B, Thaler M and Bernkop-Schnürch A, Chitosan-thioglycolic acid conjugate: an alternative carrier for oral non-viral gene delivery? J Biomed Mat Res Part A, 82(2007), 1–9. Schmitz T, Loretz B and Bernkop-Schnürch A, Development and in vitro evaluation of a thiomer-based nanoparticulate gene delivery system, Biomat, 28(2007), 524– 531. Loretz B, Thaler M and Bernkop-Schnürch A, Role of sulfhydryl groups in transfection? A case study with chitosan-N-acetyl-cysteine nanoparticles, Bioconjug Chem, 18(4)(2007), 1028–1035. Bernkop-Schnürch A, Guggi D and Pinter Y, Thiolated chitosans: development and in vitro evaluation of a mucoadhesive, permeation enhancing oral drug delivery system, J Control Release, 94(1)(2004), 177–186. Guggi D, Krauland A H and Bernkop-Schnürch A, Systemic peptide delivery via the stomach: in vivo evaluation of an oral dosage form for salmon calcitonin, J Control Release, 92(2003), 125–135. Bernkop-Schnürch A, Telsnig J and Hornof M, Development and in vitro evaluation of a mucoadhesive oral delivery system for antisense oligonucleotides, Sci Pharm, 71(2003), 165–177. Kast C E, Valenta C, Leopold M and Bernkop-Schnürch A, Design and in vitro evaluation of a novel bioadhesive vaginal drug delivery system for clotrimazole, J Control Rel, 81(2002), 347–354. Ponchel G, Montisci M-J, Dembri A, Durrer C and Duchene D, Mucoadhesion of colloidal particulate systems in the gastro-intestinal tract, Eur J Pharm Biopharm, 44(1997), 25–31. Coupe A J, Davis S S and Wilding I R, Variation in gastrointestinal transit of pharmaceutical dosage forms in healthy subjects, Pharm Res, 8(1991), 360–364. Ko J A, Park H J, Park Y S, Hwang S J and Park J B, Chitosan microparticle preparation for controlled drug release by response surface methodology, J Microencapsul, 20(2003), 791–797. Bernkop-Schnürch A, Heinrich A and Greimel A, Development of a novel method for the preparation of submicron particles based on thiolated chitosan, Eur J Pharm Biopharm, 63(2006), 166–172. Bernkop-Schnürch A, Weithaler A, Albrecht K and Greimel A, Thiomers: Preparation and in vitro evaluation of a mucoadhesive nanoparticulate drug delivery system, Int J Pharm, 317(2006), 76–81. Bernkop-Schnürch A, Kast C E and Guggi D, Permeation enhancing polymers in oral delivery of hydrophilic macromolecules: Thiomer/GSH systems, J Control Release, 93(2003), 95–103. Krauland A, Guggi D and Bernkop-Schnürch A, Oral insulin delivery: The potential of thiolated chitosan-insulin tablets on non-diabetic rats, J Control Rel, 95(2004), 547–555.

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61 Krauland A H, Leitner V M, Grabovac V and Bernkop-Schnürch A, In vivo evaluation of a nasal insulin delivery system based on thiolated chitosan, J Pharm Sci, 95(2006), 2463–2472. 62 Föger F, Hoyer H, Kafedjiiski K, Thaurer M and Bernkop-Schnürch A, In vivo comparison of various polymeric and low molecular mass inhibitors of intestinal P-glycoprotein, Biomaterials, 27(2006), 5855–5860.

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25 Chitosan–polysaccharide blended nanoparticles for controlled drug delivery J. M. A L O N S O and F. M. G O Y C O O L E A, Universidad de Santiagó de Compostela, Spain, and I. H I G U E R A - C I A P A R A, Centro de Investigación en Alimentación y Desarrollo, Mexico

25.1

Introduction

Since the early works introducing the concept of nanoparticles as colloidal drug delivery systems, much progress has taken place over the past three decades and their introduction has revolutionized the field (Marty et al., 1978; Couvreur et al., 1995; McClean et al., 1998; Soppimath et al., 2001; Panyam and Labhasetwar, 2003). The advantages of nanoparticulate carrier systems in the administration of drugs that pose difficulties associated with toxicity, degradation or poor absorption (Pinto Reis et al., 2006), along with the possibility to deliver drug molecules directly to cells in cancer therapy (Brigger et al., 2002), are presently fully advocated. In this context, polysaccharide-based nanoparticles have been a focus of increasing attention in the design and engineering of novel nanoparticulate drug delivery systems, due to their desirable properties such as biocompatibility, biodegradability, bio– and mucoadhesivity, and hydrophilic character that facilitate the administration and increase the bioavailability of poorly absorbable drugs across various epithelial barriers, such as corneal, nasal and intestinal mucosa (Alonso et al., 2007). An additional advantage of this type of system is that they can be produced under aqueous and fairly mild conditions, thus effectively, being especially suitable to preserve the bioactive conformation of delicate macromolecules (e.g. hormones, antigens, DNA, siRNA, growth factors), that otherwise would be prone to enzymatic degradation and hydrolysis (Janes et al., 2001; Csaba et al., 2006). Polysaccharide-based nanostructured drug delivery systems including chitosan–beta–cyclodextrin nanoparticles and nanocore-coated type capsules have been the focus of a recently published review by our group (Alonso et al., 2007). Also chitosan–glucomannan and chitosan–hyaluronic acid were preliminarily dealt with. In the first section of this chapter, the general properties of chitosan and other polysaccharides that have been used in combination with chitosan in the development of novel nanoparticle systems such as alginate, hyaluronic acid and konjac glucomannan, are described. This is 644 © 2008, Woodhead Publishing Limited

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followed by two sections addressing more specifically nanoparticles comprised by chitosan alone and nano-co-particle systems made of chitosan and alginate, hyaluronic acid and konjac glucomannan. For each of these systems, the methodologies and critical production parameters along with the biopharmaceutical properties and drug delivery applications are described. In the final part, an overview of the future trends and additional sources of information on the subject matter is presented.

25.2

Polysaccharides in nanoparticle formation

Polysaccharides are the most versatile kind of natural polymer due to the wide spectrum of chemical, physical and functional properties they convene in living organisms including structural, energy reserve and other more specialized roles (e.g. non self-recognition immune response in humans, signaling and resistance processes in plants, biofilm formation in microorganisms, etc.). Moreover, most polysaccharides are cheap to produce and represent by far the most abundant renewable resource available in biomass (plants, seaweeds, fungi, animal wastes and microorganisms). Despite the large tonnage of polysaccharides that are used as bulk materials in the film, fibre, textile, paper and food industries (e.g. celluloses, starch), there is increasing evidence suggesting that some polysaccharides have superior special functional properties that can be exploited in other more specialized industrial sectors, particularly in the biomedical, cosmetic and pharmaceutical fields. Especially important to the latter is the fact that some polysaccharides combine physicochemical (e.g. capacity to form gels, micro– and nanoparticles in well defined conditions) with biological properties that are key to the engineering of materials for healthcare applications. In this section we address the main properties of chitosan, alginate, hyaluronic acid and konjac glucomannan that to our view, hold promising potential in the development of a new generation of colloidal nanoparticles for transmucosal drug delivery applications.

25.2.1 Chitosan Chitosan is an aminopolysaccharide, obtained at industrial scale by thermoalkaline N–deacetylation of chitin isolated from crustacean waste. Chitosan has thoroughly been utilized in the development of potentially innovative drug delivery, tissue engineering and wound dressing systems over the last decade. Indeed, over 50% of the total number of filed patents in 2006 that claim the use of this biopolymer as a substantial part of the invention are related to drug delivery, tissue engineering and wound healing (HigueraCiapara et al., 2007). Currently, these niches are among the fastest growing markets in the world (Orive et al., 2003). This is not surprising, if we consider

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the bioactive and physicochemical properties convened by this polysaccharide that single it out from other biopolymer families. Chitosan defines a family of linear heteropolysaccharides that are comprised by (1→4)-linked 2-amino-2-deoxy-β-D-glucose (D-units) and 2-acetamido2-deoxy-β-D-glucose (A units) (Figure 25.1), obtained by deacetylation of chitin. Units of type A are often present in lower proportions than D ones, and their content, given by the molar ratio of A groups to the total (A+D), is regarded as the degree of acetylation, which is expressed either as a percentage (DA) or as a fraction (FA). Along with the degree of polymerization, the DA is a fundamental parameter that directly determines the physicochemical and biological properties of chitosan (Schipper et al., 1996; Prasitsilp et al., 2000; Huang et al., 2004). In addition to the net molar proportion of A groups, their distribution in the chitosan chain varies with the preparation protocol. Homogeneously deacetylated chitosan samples, with FA varying in the range 0.04–0.49, showed that 1H NMR diad frequencies distribution was close to a random (Bernoullian) diad distribution. In turn, chitosan samples within the same range of DA produced under heterogeneous conditions seemed to have a slightly more blockwise distribution (Vårum et al., 1991). From the physicochemical standpoint, chitosan is a water soluble polymer that can be formed into films, hydrogels, scaffolds, fibres, micro– and nanoparticles, under mild acidic conditions. Moreover, the polycationic character confers upon chitosan a high affinity for the association and delivery of therapeutic hydrophilic macromolecules (e.g. proteins, hormones, pDNA, siRNA, antigens, heparin, etc.), thus effectively, protecting their bioactivity against enzymatic and hydrolytic degradation. The physicochemical, biomedical and pharmaceutical properties of chitosan have been described in detail in several recent review articles (Skaugrud et al., 1999; Agnihotri et al., 2004; Ravi Kumar et al., 2004; George and Abraham 2006; Rinaudo 2006, 2008). A H OH

H H

O

HO

H H O

H NH2 H

H

H O O

O

HO

H OH

NH2 H

O

HO

H H

H OH

NH

H

CH D

D

O

CH3

25.1 Schematic representative structure of a chitosan chain comprising β(1→4)–linked 2–acetamido–2–deoxy–β–D–glucopyranose (A–residues) and 2–amino–2–deoxy–β–D–glucopyranose (D–residues) monosaccharide residues.

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The biological properties of chitosan in animals have exhaustively been studied over the past decades and we will only discuss them briefly here. Chitosan has been widely used in the food industry and is an approved food additive in Japan. It is marketed in Europe, USA and many more countries as a fat binder in cholesterol-lowering and slimming formulations (Shahidi et al., 1993). A monograph relating to chitosan hydrochloride for biomedical use was included in the fourth edition of the European Pharmacopoeia (EDQH, 2002). Chitosan is known to be biodegraded by several enzymes, among them chitinases secreted by intestinal microorganisms and lysozyme, which is highly concentrated in mucosal surfaces (Pangburn et al., 1982; Aiba, 1992; Hirano et al., 1990). Other biological properties have thoroughly been investigated in chitosan, both in vitro and in animal models. Mucoadhesivity of chitosan and some derivatives has been recognized (Lehr et al., 1992; Shruti et al., 2006). Chitosan is also well known to be biocompatible (Hirano et al., 1990; Chatelet et al., 2001; Vandevord et al., 2002), non toxic (Aspden et al., 1997), immunostimulatory (Peluso et al., 1994; Babensee and Paranjpe, 2005; Porporatto et al., 2005; Borges et al., 2007), and to exhibit capacity to promote the absorption of poorly absorbable macromolecules across epithelial barriers by transient widening of cell tight junctions (Artursson et al., 1994; Kotze et al., 1997; Schipper et al., 1997; Jung et al., 2000; Smith et al., 2004).

25.2.2 Alginate Alginate is yet another polysaccharide that has drawn substantial attention in the biomedical field over the past decades and is the subject matter of Chapter 19 of this book, therefore we will only focus here on some of the key properties relevant to the current applications in nanoparticle-mediated drug delivery. Chemically, alginate is a linear block co-polymer constituted by homopolymeric poly-L-guluronate (poly-G) and poly-D-mannuronate (polyM) and by alternate blocks of both residues. Its composition, summarized by the molar ratio of D-mannuroate to L-guluronate residues (M/G ratio), is the main property that dictates the physicochemical and functional properties of alginate: Namely, the gelling capacity in the presence of specific bivalent cations (Ca2+, Ba2+ and Sr2+ and not Mg2+) as a result of a cooperative ‘eggbox’ type of interaction between poly-G blocks of more than 6–10 residues, where every two carboxylate groups form a coordination complex with the cation (Grant et al., 1973). As in the case of chitosan, the biological properties of alginate which are relevant for its application in drug delivery can be summarized in three main categories: biocompatibility, bioadhesivity and immunogenic activity. Alginates are included in a group of compounds that are regarded as generally safe (GRAS) by the FDA. The oral administration of alginate has not been shown

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to be toxic or to provoke much immunogenic responses (Espevik et al., 1993). Commercial alginates when tested after purification by free-flow electrophoresis did not provoke foreign body reactions at least three weeks after implantation in the peritoneal cavity of rodents (Mumper et al., 1994). Other studies have reached similar conclusions from in vitro mice lymphocyte tests as well as by implantation of Ba2+ crosslinked beads beneath the kidney capsule of BB/OK rats (Klöck et al., 1997; Shapiro and Cohen, 1997; Jork et al., 2000; Orive et al., 2005). Derived from their biocompatibility along with their gelling capacity, alginates have been applied in the engineering of biomaterials (Rajaonarivony et al., 1993; Klöck et al., 1994; Shapiro and Cohen, 1997; Jork et al., 2000; Augst et al., 2006). It has also been known for a long time that alginate exhibits bioadhesive properties (Chen and Cyr, 1970; Smart et al., 1984). However, it is only recently that the potential benefit of this property began to be explored. Particularly, bioadhesion of alginate to oesophagus tissue has been exploited to develop a treatment for upper gastro-oesophageal disorders including gastro-oesophageal reflux disease (GORD) where coating with liquid alginate formulations exerts a protective effect (Batchelor et al., 2002). More recently, it was shown that alginate can protect against the effects of bile acids on the onset of Barrett’s esophagus, a metaplastic condition and precursor to oesophageal adenocarcinoma (Dettmar et al., 2007). This is certainly an area of great opportunity for the application of alginate-based nanoparticles for drug delivery aimed at oesophagus disorder therapy. Finally, with regard to the immunostimulatory activity of alginate the results published until now are not yet conclusive as conflicting evidence stems from differences in M/G ratio, Mw and purity among alginate samples used for different studies (Babensee and Paranjpe, 2005; Kurachi et al., 2005; Orive et al., 2005; Borges et al., 2007).

25.2.3 Hyaluronic acid Along with chitosan and alginate, hyaluronic acid is the other main polysaccharide considered as a good basis for biomedical applications, mainly due to its recognized biological activity (Milas and Rinaudo, 2004). Hyaluronic acid (also named hyaluronan) is also an aminopolysaccharide that occurs in the extracellular matrix of connective tissues in vertebrates, such as subcutaneous tissue, synovial fluid, cartilage, umbilical cord and vitreous and aqueous humor in the eye. In the past, hyaluronan used to be sourced from animal waste of bovine vitreous humor, roster combs or umbilical cords; then, it was very expensive and usually had undesirable bound proteins. The same polysaccharide is now successfully produced on a large scale by the bacteria Streptococcus zooepidemicus and Streptococcus equi with a good yield and high degree of purity (Milas and Rinaudo, 2004). Thus, the price has decreased, allowing the development of many new applications. A

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couple of recent reviews account for the main biomedical applications of hyaluronic acid (Kogan et al., 2007; Rinaudo, 2008). The chemical structure of hyaluronic acid is described as a linear (AB) charged copolymer based on β(1→4)-linked D-glucuronic acid (GlcA) and β(1→3)-linked N-acetyl-Dglucosamine (GlcNAc) alternated monomeric units (Figure 25.2). Hyaluronic acid is considered not only a pharmaceutical excipient but also a drug. Its viscoelasticity and biocompatibility has led to its utilization in ophthalmology, dermatology and osteoarthritis. From the biopharmaceutical stand point, the main interest in this polysaccharide is in its bio– and mucoadhesive properties, particularly because of the specific hyaluronan (CD44) receptors known to be present in biological barriers (Ahlo and Underhill, 1989; Knudson et al., 2002).

25.2.4 Konjac glucomannan Konjac glucomannan is yet another food-grade polysaccharide, sourced from the tubers of Amorphophalus konjac, an edible plant crop in Asia that has recently been found to be appealing for the development of novel chitosanbased nanoparticles for drug delivery applications (Alonso-Sande et al., 2008a). By contrast with chitosan, alginate and hyaluronic acid, which bear charged groups in their structure, native konjac glucomannan is a neutral polysaccharide having a backbone chain comprised by β(1→4)-linked D-glucose and Dmannose in a ratio of ~1.6:1 with branching occurring at D-mannose residues through 1→3 linkages, approximately three for every ~32 sugar residues as shown in Figure 25.3a (Maeda et al., 1980). The phosphorylated derivative of konjac glucomannan is also available commercially (Figure 25.3b). The biopharmaceutical properties that account for the behaviour of konjac glucomannan in nanoparticle drug delivery systems include its high content of mannose, thus effectively enabling the interaction of konjac glucomannan with biological surfaces which are particularly rich in mannose receptors such as M-cells overlying Peyer’s patches (Takada et al., 1984; Tomizawa et al., 1993) and macrophages (Cui et al., 2003). O H

OO–

CH3

H H O

O

H

NH

OH H

HO H

OH H

H O

H O H

OH

25.2 Schematic representative structure of hyaluronic acid comprising alternating β(1→4)–linked β–D–glucuronic acid (A–residues) and β(1→3)–linked 2–acetamido–2–deoxy–β–D– glucopyranose (B–residues) monosaccharide residues.

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H

OR′

H

HO O

O

HO

H

O

H

H

H

OH

OH O

O

O

HO

H

H OH

H

H

O R″ O

H OR′

H

H

H

H

n=11-30 H HO H O

H

HO R″

(0.05) –H (0.95)

R′

HO

CH3

H H

(a)

O

OH

O P O

O O

H H

O

H O

H HO

P H

O O

HO O

O

O

H

H

H

O HO

H

H

H OH

HO

H

n=0.07

H H HO HO

H (b)

25.3 Schematic representative structure of: (a) konjac glucomannan chain comprising β(1→4)–linked β–D–glucopyranose and β–D– mannopyranose monosaccharide residues, with sporadic β(1→4)– linked branched at β–D–mannopyranose sugar residues; and (b) representative sequence present in phosphorylated konjac glucomannan.

In order to keep the focus, we have intentionally omitted to include in this review other polysaccharides that, though amenable to interact with chitosan and harness nanometric structures, thus far, only little work has been conducted on their utilization in the development of nanocarriers for biopharmaceutical applications (e.g. dextran and dextran derivatives, carboxymethyl cellulose, chondroitin sulfate, heparin and heparin sulfate, dermatan sulfate, keratan sulfate, pectin, xanthan and gellan). To finalize this section, it is important to mention that cyclodextrins are also a very important class of oligosaccharides with firmly established use in pharmaceutical applications, that has also been incorporated into chitosan

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nanoparticle drug delivery systems as a strategy to increase the loading capacity of lipophilic drugs (Maestrelli et al., 2006) and to modulate the drug release in vitro of insulin and heparin (Krauland and Alonso 2007).

25.3

Nanoparticles constituted by chitosan

Most frequently, chitosan nanoparticles are formed according to a bottom-up approach as a result of a self-assembling or crosslinking processes in which the molecules arrange themselves into ordered nanoscale structures either by physical or covalent inter– or intramolecular interactions. In these nanostructures, the drug can be entrapped or attached to the nanoparticle matrix. Matrix nanoparticles have been prepared by several methodologies that have been summarized in Figure 25.4. Often, a combination of methods is utilized (e.g. chemical surface crosslinking of preformed nanoparticles obtained by coacervation). It is important to point out that despite the variety of methods documented to harness chitosan nanoparticle formation, only a few offer real potential for pharmaceutical applications, due to the strict regulations that excipients intended for drug delivery formulations must comply with. These methods are mostly based upon the principle of ionic gelation of chitosan. Consequently, to keep the focus, only the principles behind the formation of chitosan nanoparticles by ionic gelation are briefly discussed next. Polymer matrix nanoparticles that are formed as a consequence of a solto-gel process such as in ionotropic gelation, are governed by essentially the same general mechanisms as those underlying the creation of macroscopic polymer gels under quiescent conditions, leading to a homogeneous percolated three-dimensional gel network. However, by contrast with ‘bulk’ gel systems, nanoparticle gelation occurs in conditions that avoid macrogelation. This is achieved provided the polymer concentration is below the critical gelling concentration (Co) and by controlling the solvent solubility parameters. Colloidal stability of gel nanoparticles against their further aggregation and growth with time is achieved either by electrostatic charge repulsion at their surface, characterized by a high zeta potential or by steric stabilization by the presence of bulk hydrophilic chain ‘loops’ and dead ends at the surface (e.g. PEO–PPO diblock copolymer) (Calvo et al., 1997b). Colloidal stability results from the balance between attractive and repulsive forces on the surfaces of the particles in a given medium (i.e. water), governed by electrostatic and van de Waals interactions. This is achieved when the Hamaker constants of the interacting particles and water are matched (Vinogradov et al., 2002). In general, chitosan gelation can be achieved either by physical or covalent crosslinking approaches. Physical methods of gelation include, among others, ion bridging by phosphorous-containing ions such as pentasodium tripolyphosphate (TPP) (Shiraishi et al., 1993; Remuñán-López and Bodmeier,

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Ionic crosslinking

Ionotropic gelation (e.g. tripolyphosphate)

Chemical crosslinking

Glutaraldehyde

Ethylene diglycidyl ester (EDGE)

25.4 Methods for preparing chitosan nanoparticles.

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Coacervation

Genipin

Precipitation

Self-assembly (hydrophobic derivatives)

Complex coacervation

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Chitosan nanoparticles

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1997; Calvo et al., 1997a; Zhang et al., 2004), polyphosphoric acid (Mi et al., 1999b, c; Lee et al., 2001); hexametaphosphate (Gupta and Jabrail, 2006), hydrophobic association induced by charge neutralization with betaglycerol phosphate (Chenite et al., 2001); as well as selective ion interactions with molybdate polyoxoanions (Draget et al., 1992; Dambies et al., 2001), Pt(II) (Brack et al., 1997), EDTA (Valenta et al., 1998), and by interaction with anionic surfactants such as sodium dodecyl sulphate (SDS) (Babak et al., 2000). Despite the numerous documented substances found to induce the ‘bulk’ physical gelation of chitosan, only a few of them have exhibited no physiological toxicity and can bear any real potential for commercial application in the pharmaceutical field. Among these, chitosan–TPP (Calvo et al., 1997a) and chitosan–EDTA (Loretz and Bernkop-Schnürch 2006) have been exploited in the formation of nanoparticle systems for drug and gene delivery. Our research group pioneered the development of a nanoparticle drug delivery platform based on a new type of chitosan nanoparticle obtained by ionotropic gelation of chitosan with pentasodium tripolyphosphate (TPP) (Calvo et al., 1997a). Among the main advantages of this hydrophilic system is that it avoids the use of high temperatures, organic solvents or sonication. Such mild conditions along with the polycationic nature of chitosan confer on this nanoparticle system the capacity to effectively associate and preserve the stability and bioactivity of therapeutic macromolecules against enzymatic and hydrolytic degradation. Indeed, this system has shown high loading capacity for hydrophilic proteins such as insulin (Fernández-Urrususno et al., 1999; Pan et al., 2002; Ma et al., 2002, 2005; Vila et al., 2004), bovine serum albumin (Xu and Du, 2003; Zhang et al., 2004) and tetanus toxoid (Vila et al., 2004). The protein loading capacity of these systems has been found to be mostly affected by the nature of the protein (e.g. solubility and pI) (Vila et al., 2002) and by the theoretical protein load itself (Xu and Du, 2003), pH (Ma et al., 2002), as well as by the molecular weight (Mw) of chitosan (Vila et al., 2004). Other therapeutic macromolecules loaded into these nanoparticles include: pDNA (Mao et al., 2001), heparin (Krauland and Alonso, 2007), hydrophobic peptides such as cyclosporine A (De Campos et al., 2001; El-Shabouri, 2002) and smaller molecules such as ammonium glycyrrhizinate (an anti-inflammatory drug) (Wu et al., 2005). The interaction of chitosan with TPP is accepted to be mediated by the 5– 4– intramolecular crosslinking of tripolyphosphoric (P3 O 10 and HP3 O 10 ) ionic species, products of the dissociation of TPP in aqueous solution, with –NH 3+ groups in chitosan. The role of the pH of the TPP solution in the crosslinking mechanism has been carefully accounted for by use of FTIR spectroscopy 5– ions and potentiometric methods (Mi et al., 1999a). At pH 4.0, only P3 O 10 exist and the crosslinking occurs only via the ionic crosslinking between these ions and –NH 3+ . In the original solution of TPP (pH ~9.7) however, the 5– 4– concentration of P3 O 10 and HP3 O 10 is still high but there are also OH– ions

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that may also play a role in neutralizing –NH 3+ charged groups in chitosan. As a consequence, the particle size in this system is expected to be highly sensitive to the pH. In line with this, other studies have revealed that CS– TPP nanoparticles experience volume phase transitions induced independently by variations in pH and ionic strength (López-León et al., 2005). The effect of other preparative conditions and intrinsic characteristics of chitosan on the physical characteristics of chitosan–TPP nanoparticles has been studied, with attention to the ratio of chitosan to TPP (Xu and Du, 2003; Zhang et al., 2004; Gan et al., 2005); chitosan Mw (Janes et al., 2003, Vila et al., 2004; Goycoolea et al., 2007) and DA (Xu and Du, 2003; Goycoolea et al., 2007). Several features can generally be recognized in chitosan–TPP nanoparticles obtained under various preparation conditions using chitosan samples of different Mw and DA: the average diameter, as determined by dynamic light scattering, reported for a large series of nanoparticle samples, varies from 100 to 354 nm; a positive zeta potential lying in the range +20 to +48 mV; a spherical morphology as visualized under transmission electron (TEM) and also by atomic force microscopy (AFM) (Kim and Kang, 2006). It is important to point out that the hybrid nanoparticle systems comprising ionotropically gelled chitosan and other polysaccharides (e.g. alginate, hyaluronic acid, konjac glucomannan), differ from systems obtained by simple polyelectrolyte complexation between chitosan and an oppositely charged polyanion (e.g. DNA, proteins, polyanionic polysaccharides) in the following aspects: (a) nanoparticles have a more defined homogeneous structure; (b) they can co-entrap different ingredients; and (c) DNA, proteins or anionic polysaccharides are not displaced from the nanoparticles by the presence of a competitor. Addressing the properties and applications of drug and gene delivery systems derived from chitosan-based polyionic complexes lies beyond the scope of this chapter.

25.4

Drug delivery properties and biopharmaceutical applications

As a result of its rather unique properties, chitosan has been utilized in pharmaceutical research to foster the development of novel nanoparticle vehicles. A recognized feature of these nanosystems is their capacity to protect macromolecules against degradation and their ability to overcome mucosal barriers. As a consequence, their application has been mainly centred in non-invasive routes of administration i.e. ocular, nasal and oral mucosa (Janes et al., 2001; Alonso and Sanchez, 2003, 2004; Prego et al., 2005; Sanchez and Alonso, 2006; Csaba et al., 2008). Also, recent studies have evidenced the feasibility to include chitosan–TPP nanoparticles into spraydried mannitol-based microparticle systems for administration via the pulmonary epithelium (Grenha et al., 2005, 2007). A more detailed explanation

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of the potential of these systems for different mucosal modalities of administration is described below.

25.4.1 Ocular administration The efficacy of chitosan nanoparticles to prolong the delivery of drugs to the eye surface has been firmly demonstrated (Alonso and Sanchez, 2003; Sanchez and Alonso, 2006). This has been achieved using cyclosporine A (CyA) as a model of a drug that could benefit from a prolonged delivery to the eye surface due to its potential indication for the treatment of severe dry eye. Interestingly, chitosan nanoparticles were able to provide a selective and prolonged delivery of CyA to the ocular mucosa without compromising inner ocular tissues and avoiding systemic absorption (De Campos et al., 2001). It was also found that chitosan nanoparticles are able to provide a selective and prolonged delivery to the ocular mucosa, avoiding systemic absorption. In subsequent studies (De Campos et al., 2004), using fluorescently-labelled chitosan it was possible to show that chitosan nanoparticles adhere to the cornea and conjunctiva and remain associated to them for more than 24 h.

25.4.2 Nasal administration Chitosan nanoparticles have been shown to improve the administration of therapeutic and antigenic macromolecules across the nasal mucosa as has recently been reviewed (Köping–Höggård et al., 2005; Csaba et al., 2008). The studies on the efficacy of CS nanoparticles for improving the nasal delivery of macromolecules have been performed with two types of molecules: the polypeptide insulin and the antigen tetanus toxoid (Fernández-Urrusuno et al., 1999; Vila et al., 2002, 2004). With respect to insulin, we have evaluated the efficacy of chitosan nanoparticles at increasing the hypoglycaemic response in either normal rabbits or diabetic rats. The results obtained in normal rabbits indicated that chitosan nanoparticles led to a significant decrease of the glycaemia levels (40% maximum reduction) compared to an insulin solution containing chitosan (Fernández-Urrusuno et al., 1999). We have also studied the systemic and mucosal immune response after nasal administration of antigen-loaded chitosan nanoparticles to conscious mice (Vila et al., 2004). It was found that chitosan nanoparticles are able to elicit high and long-lasting humoral as well as mucosal immune responses, as evidenced from IgG levels and virtually no differences were observed in the antibody patterns of nanoparticles made with 23 kDa chitosan or 70 kDa chitosan. This increasing and long-lasting response has led us to hypothesize that the nanoparticles cross the nasal mucosa and reach the antigen presenting cells. In this particular environment, the particles might deliver the associated antigen for extended periods of time. The mechanisms whereby chitosan

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nanoparticles improve the nasal transport of macromolecules are understood on the basis of their facilitated interaction with the nasal epithelium.

25.4.3 Oral administration Chitosan-based nanoparticles have also been extensively studied as vehicles for oral administration of macromolecular drugs (Prego et al., 2005). Insulinloaded chitosan nanoparticles prepared by ionotropic gelation have been administered to diabetic rats (Pan et al., 2002). The reduction in the glucose plasmatic levels gave evidence of the capacity of the nanoparticles to enhance the intestinal absorption of the peptide. This effect was prolonged for 15 hours and the pharmacological bioavailability was 14.9%. Two further in vivo studies conducted in diabetic rats confirmed these effects with somewhat lower pharmacological bioavailability values (Cui et al., 2004; Ma et al., 2005). The experimental evidence available on the oral administration of peptides in chitosan-based nanoparticles suggests that the nanoparticles are able to firmly adhere and enter the epithelia and provide a continuous delivery of the peptide to the blood stream. This hypothesis was corroborated by the observed internalization of the nanosystems in the enterocyte-like Caco-2 monolayer cell line and with mucus producing MTX–E12 cell membranes. This correlated well with their uptake in vivo in a rat model (Behrens et al., 2002). Results indicated that chitosan nanoparticles associated more strongly to MTX–E12 than to the more hydrophobic polystyrene systems. An intense electrostatic interaction between positively charged chitosan nanoparticles and negatively charged mucins is regarded as the cause for the strong interaction of chitosan with mucus. Indeed, removal of the mucus layer led to a substantial decrease in the association of chitosan nanoparticles to MTX–E12. However, confocal laser scanning microscopy (CLSM) studies revealed that chitosan also associated strongly with Caco-2 cell monolayers (0.056 ± 0.004 µg/ mm2, 7.8%) by an electrostatic interaction with negatively charged glycocalyx on the apical membrane. The in vitro results were validated by uptake studies by assessing the interaction of nanoparticles with different sections of small intestine after intra–duodenal injection in rats. A significant amount of nanoparticles was taken up by both ephitellial cells and Peyer’s patches. The transport of chitosan nanoparticles through Caco-2 cells was found to be controlled by endo-and transcytosis.

25.5

Hybrid nanoparticles consisting of chitosan and other polysaccharides

Chitosan-based nanoparticles have been prepared by incorporating other polysaccharides, including alginate, konjac glucomannan, hyaluronic acid, as well as oligosaccharides such as cyclodextrins. Table 25.1 summarizes a

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Table 25.1 Examples of chitosan–based nanoparticle systems and proof–of–concept studies for their use as transmucosal carriers of bioactive macromolecules Other components

Macromolecule

Size (nm)

Zeta potential (mV)

Assoc. efficiency (%)

Admin. route

Animal model

Ref.a

Protasan® 110 CL DA 13%; Mw < 5 × 104 Da

TPP (CS:TPP mass ratio 6:1)

Insulin

352±11

+32.9±0.2

92.5±0.5

Nasal

Rabbit

[1]

Aldrich® DA 15%; Mw ~ 1.8 × 105 Da

TPP (CS:TPP mass ratio 4:1)

Insulin

269±7

+34.9±0.7

38.5±1.5

Oral

Diabetic rat

[2]

Protasan® DA 13%; Mw > 2 · 3 × 104 Da

TPP (CS:TPP mass ratio 6:1)

Tetanous toxoid

354±27

+36.9±0.3

55.1±3.4

Nasal

Mice

[3]

Sea Cure® Cl213 DA 16%; Mw ~2.7 × 105 Da

TPP (CS:TPP mass ratio 8:1)

Cyclosporine A

239±9

+37.5±0.9

73.4

Ocular

Rabbit

[4]

Protasan® DA 13%; Mw ~ 1.1 × 105 Da

TPP and alginate (CS:TPP:alginate mass ratio 6:1:0.5)

Insulin

275±1

+42.8±4.0

55.8±1.2

Nasal

Rabbit

[5]

Protasan® DA 13%; Mw ~ 1.1 × 105 Da

TPP and hyaluronate (CS:TPP:hyaluronate mass ratio 6:0.3:3)

Plasmid (pEGFP-C1)b

180±36

+20±0.6

High

Ocular

Rabbit

[6]

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Notes: a References cited in Table: [1] Fernandez-Urrusuno et al. (1999) [2] Ma et al. (2005) [3] Vila et al. (2004) [4] De Campos et al. (2001) [5] Goycoolea et al (2007b) [6] De la Fuente et al. (2008c) b pDNA encoding for green fluorescent protein (pEGFP-C1)

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non-exclusive list of examples of chitosan-based nanoparticles systems and their applications as carriers for bioactive molecules. Due to the number of new studies, we will address below in detail chitosan–alginate, chitosan– konjac glucomannan and chitosan–hyaluronic acid systems.

25.5.1 Chitosan–alginate nanoparticles Methodologies and critical production parameters Three distinctive strategies to prepare chitosan–alginate co-gelled nanoparticle systems and their utilization as carriers of therapeutic macromolecules have been described. They are represented schematically in Figure 25.5 and their main physicochemical and pharmacological properties are described next. Chitosan-coated ionotropically pre-gelled alginate nanoparticles A two-step procedure has been described by virtue of which alginate is pregelled in the presence of calcium chloride under sonication and then the nanogels are further coated with chitosan (De and Robinson, 2003). To this end, a method first published by Rajaonarivony et al., (1993) was modified. Calcium chloride at about 7% of ‘egg-box’ stoichiometry is added to an aqueous dilute alginate solution (~0.6 mg/mL) under sonication, leading to the formation of pre-gelled primary nanoparticles. A chitosan solution (in acetic acid 3%) is then added to the alginate nanoparticle suspension. The alginate–chitosan nanoparticles thus obtained had an average size of 650 ± 22 nm provided the calcium chloride to sodium alginate mass ratio is at the

Alginate matrix ionically crosslinked with CaCl2

Chitosan coat

Chitosan matrix coacervated with Na2SO4

Alginate coat

Alginate

TPP Chitosan (a)

(b)

(c)

25.5 Schematic representation of types of chitosan–alginate nanoparticle systems: (a) Alginate nanomatrix prepared by ionotropic gelation with calcium chloride coated with chitosan; (b) Chitosan nanomatrix prepared by coacervation with sodium sulfate coated with alginate; (c) Chitosan–alginate co-gelled hybrid nanoparticle by ionotropic gelation of chitosan with pentasodium tripolyphosphate (TPP) and concomitant complexation of alginate.

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pre-gel condition, in the range from 0.08 to 0.35 (optimal at ~0.1) at a constant chitosan to sodium alginate ratio of ~0.1. Beyond such a calcium chloride to alginate ratio, particles were found to grow to large aggregates. The role of the main parameters governing the physical characteristics of this system has also been studied. For example, Douglas and Tabrizian (2005), evaluated the effect of chitosan (DA 15%; Mw 1000 and 50 kDa) and alginate (Mw ~ 120–190 kDa, 80–120 kDa and 12–80 KDa) Mw, pH and alginate to chitosan mass ratio on the particle size. The results showed that, in general, the use of the low Mw polymers resulted in smaller particles for most alginate to chitosan mass ratios. The smaller particles (~338–415 nm) were attained at alginate to chitosan mass ratios close to ~1:1, a contrasting result to De and Robinson’s (2003) who found the smaller size at ratios of ~30:1. Possible discrepancies were attributed to differences in pH. Alginate-coated chitosan nanoparticles by coacervation A three-step protocol has been described to obtain protein-loaded nanoparticles as an optimized delivery platform for antigens (Borges et al., 2005). To this end, chitosan nanoparticles are initially preformed by mixing sodium sulfate solution to a chitosan solution in the presence of a surfactant (Tween® 80), under mild agitation and continuous sonication, according to an early protocol (Berthold et al., 1996). The isolated particles are then freeze-dried, resuspended in phosphate buffer (pH 7.4) and incubated in solutions of ovalbumin of varying concentration. Ovalbumin-loaded particle suspensions were further added to alginate solutions of varying concentration and lastly, the alginate at the surface was cross-linked with calcium chloride. Particle size remained in the range of 100 to 1000 nm (average 643.2 ± 171.7 nm); however, large polydispersity values were invariably found for this system, attributed to the freeze-drying process. Ovalbumin association efficiency and loading capacity achieved in this system were 40 and 80%, respectively. The formation of an alginate coating layer was confirmed by the inversion of the zeta potential from +41 mV for the uncoated nanoparticles to –35 mV after the coating step, as was expected from the contribution of the alginate negative charge. This was also confirmed by changes in the main spectral FTIR bands corresponding to chitosan N–H (1542 cm–1) and C–N stretch (1414 cm–1) and alginate carboxylate asymmetric stretch (1605 cm–1). The alginate coating was found to stabilize the nanoparticles in SIF buffer (pH 6.8) at 37°C, and, as a consequence, enabled the prevention of a burst release of loaded ovalbumin. The optimization of the coating process resulted in 35% (w/w) for the loading capacity of the coated particles. In addition, the SDS–PAGE analysis indicated that the structural integrity of ovalbumin released from the nanoparticles was unaltered.

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Chitosan–alginate hybrid co-gelled nanoparticles We have recently developed a new type of hybrid nanoparticle system, made by ionic gelation of chitosan with TPP and concomitant complexation with alginate (Goycoolea et al., 2008). To this end, alginate with Mw in the range 4–32 KDa was incorporated in the system in a chitosan/alginate mass ratio ~10:1 and the physicochemical properties of the system were evaluated in terms of size, zeta potential, morphology, association of insulin and stability in acetate buffer (pH 4.3 at 37°C) in a system intended for nasal administration. Spherical nanoparticles were formed as characterized by TEM, with a hydrodynamic diameter in the range 275–389 nm, low polydispersity index (PI < 0.36) and zeta potential +43.1 ± 4.0 mV. These nanoparticles remained stable in acetate buffer for up to 80 min. Insulin was associated with efficiencies of ~50% and more than ~80% was released in acetate buffer (pH 4.3 at 37°C) within the first 20 min regardless of the Mw of alginate.

25.5.2 Drug delivery and biopharmaceutical applications Methylene blue has been used as a model molecule to investigate the drug release characteristics of the chitosan-coated ionotropically pre-gelled alginate nanoparticles developed by De and Robinson (2003). It was found that only 2% of methylene blue was released after 135 hours in the absence of added sodium chloride. However, upon sequential incremental additions of this electrolyte, the cumulative amount of released methylene blue was directly proportional to the concentration of added electrolyte and linear dependence was established between these two parameters as shown in Figure 25.6. This was explained as a consequence of the strong electrolyte character of sodium chloride leading to calcium–sodium ion exchange that leads to solubilization of calcium alginate and to a destabilization of the nanoparticle surface by competing with chitosan. This chitosan–alginate system has also been applied to the oral delivery of insulin (Sarmento et al., 2006). Up to 91% of insulin association efficiency was achieved for some formulations. The in vitro release properties of these nanoparticles were evaluated in simulated gastric and intestinal fluids. The results of these studies showed the protective role of chitosan against insulin diffusion from the alginate matrix due to the formation of an alginate/chitosan complex. In addition, the preservation of the secondary structure of insulin entrapped within this system was also probed by FTIR spectroscopy and circular dichroism (CD) (Sarmento et al., 2007). FTIR second derivative spectra and area overlap compared to an insulin standard confirmed that no significant conformational changes of insulin occurred in terms of α-helix and β-sheet content. Far-UV–CD spectra corroborated the maintenance of the insulin structure during the nanoparticle production.

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80 Initial conc. of 0 mg/mL Initial conc. of 1 mg/mL Initial conc. of 3 mg/mL Initial conc. of 6 mg/mL

70

% Cum. MB release

60 50 40 30 20 10 0 0

20

40

60

80 Time (hr)

100

120

140

25.6 The release of methylene blue from chitosan–alginate nanoparticles as a function of time and sodium chloride concentration. Arrow (↑) identifies the time points where incremental amounts of NaCl were added to the media. Source: De and Robinson (2003). With permission of Elsevier Scientific.

In a further study, the uptake of these alginate-coated nanoparticles into rat Peyer’s patches though M-cells was investigated by labelling the ovalbumin with a fluorescent tag (Borges et al., 2006). The nanoparticles were internalized by rat Peyer’s patches. Fluorescent nanoparticles were visualized not only in the region underneath (0.5 µm) of the follicle-associated epithelium, the subepithelial dome region, but also in deeper regions of the secondary lymphoid organ. It is suggested that the nanoparticles are most likely internalized by phagocytic cells, most probably by dendritic cells present in Peyer’s patches. The results of this study, compared with previous work undertaken with chitosan nanoparticles (Behrens et al., 2002), confirm the notion that cellular uptake is the result of a complex combination of size, and specific interactions of motifs on the nanoparticle surface with the cellular membrane in mediating endocytotic uptake by nonmucus covered intestinal cells, e.g. M-cells. In turn, chitosan–TPP co-gelled with alginate systems (Goycoolea et al., 2008) were evaluated as vehicles for nasal administration in rabbits. The results showed that these nanoparticles were able to facilitate the transport of insulin across the nasal mucosa, leading to a significant absorption (40% reduction in the plasma glucose levels) which was prolonged for several hours.

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25.5.3 Chitosan–hyaluronic acid systems Methodologies and critical production parameters Chitosan and hyaluronic acid have been associated in the form of hybrid cogelled nanoparticles using the same ionotropic gelation process described in previous sections with a view to promoting the utility of the nanocarrier for the transmucosal delivery of macromolecules (De la Fuente et al., 2008a). Briefly, these nanoparticles were formed by a simple mixing of a CS aqueous solution with an aqueous solution of hyaluronic acid and the gelation agent TPP. Incorporation of hyaluronic acid in this system was probed by agarose gel electrophoresis, whereas the stability in biological conditions upon incubation in simulated lachrymal fluid (SLF) was studied from the evolution of the particle size and zeta potential during 24 h. These nanoparticles exhibited a spherical shape, size in the nanometric range and a positive zeta potential. These parameters varied, as expected, depending on the chitosan/hyaluronic acid mass ratio. The results of the particle size tests confirmed that nanoparticles were formed in a hyaluronic acid/chitosan mass ratio range between 0.1:1 and 1:1 and a TPP concentration from 0.25 to 1 mg/mL. Both, hyaluronic acid and TPP concentrations affect the size of the nanoparticles. More specifically, an increase in hyaluronic acid amount was accompanied by an elevation in the nanoparticle size. This result was taken as evidence of the effective incorporation of hyaluronic acid into the nanoparticle structure and suggests that the particle formation mechanism is partially driven by electrostatic interactions between the polysaccharides, as previously documented for the formation of chitosan–hyaluronate polyelectrolyte complexes (Rusu-Balaita et al., 2003). These nanoparticles exhibited a positive surface zeta potential with values ranging between ~+30 and ~+47 mV. Hyaluronic acid concentration also had an effect on the zeta potential In general, increase in the amount of hyaluronic acid leads to a reduction in the zeta potential value, which has been interpreted as the expected consequence of preferential location of some hyaluronic acid molecules on the surface of the nanoparticles. In line with chitosan–konjac glucomannan hybrid nanoparticles, an additional technological asset of chitosan–hyaluronate nanoparticle systems is that they can also be effectively freeze-dried and resuspended by hand shaking. Drug delivery and biopharmaceutical applications Chitosan–hyaluronic acid nanoparticles have been loaded with BSA, insulin, hydrophobic peptides such as CyA, as well as with a model plasmid encoding for a green fluorescent protein. In all cases, the association efficiency is very high (close to 90% for both protein and pDNA), much the same as that found in pure chitosan nanopartices. In the case of BSA, it has been found that

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hyaluronic acid does not show an influence on the association efficiency. This has been explained as the consequence of the interaction of BSA with hyaluronic acid by virtue of hydrophobic interactions, H-bonding, and other intermolecular forces, as was observed in the preparation of hyaluronic acidBSA complexes (Xu et al., 2000). In turn, CyA has also been incorporated in this nanoparticle system, aiming to evaluate the association of a model polypeptide of notable hydrophobicity, neutral charge, lypophilic character and widely used in therapeutics as an immunosuppressant agent. This was done bearing in mind that the co-administration of CyA and hyaluronic acid intravenously leads to a fourfold increase in the immunosuppressive effect of the polypeptide. This was attributed to the interaction of hyaluronic acid to the CD44 receptor (Gowland, 1998). Moreover, as discussed earlier in this chapter, the utility of chitosan-based nanoparticle systems for improving the bioavailability of CyA after oral (El-Shabouri, 2002) and ocular (De Campos et al., 2001) administration has been demonstrated. Therefore, given the merging of biopharmaceutical advantages of hyaluronic acid and chitosanbased nanoparticles, this system appears as an excellent candidate for the transmucosal delivery of complex macromolecular compounds. Other related studies conducted by our group have aimed to develop new chitosan–hyaluronic acid nanocarrier systems for gene delivery (De la Fuente et al., 2008b). To this end, chitosan of Mw ~ 125 kD or ~10–12 kDa and hyaluronic acid of Mw ~ 170 or < 10 kDa were utilized to prepare nanoparticles by ionic crossliking using TPP. The biocompatibility and the ability of promoting gene transfection of these nanoparticles were studied at cellular level and their internalization allowed exploring the mechanisms mediating cell transfection. Two plasmids (pDNA) encoding for green fluorescent protein (pEGFP–C1) and for beta galactosidase (pβ-Gal) were associated to the nanoparticle system by mixing them with the hyaluronic acid-TPP solution prior to adding it into chitosan. Cell line HEK 293 was used to study the cytotoxicity, gene transfection efficiency and uptake and intracellular distribution studies. Gel retardation assays confirmed that more than 86% of pDNA was associated into the various formulations tested comprising chitosan and hyaluronic acid of high and low Mw. It was also found that this system preserves intact the conformation and thus the bioactivity of the plasmid after incubation with chitosanase and hyaluronidase as probed by gel electrophoresis. Enzymatic degradation of the chitosan in the polymer matrix was the major mechanism governing the release of pDNA. As regards the biopharmaceutical characteristics featured by chitosan– hyaluronic acid nanoparticle system, it is worth pointing out its low cytotoxicity irrespective of the polysaccharide’s Mw. Even at doses up to 154.4 µg of nanoparticles/cm2 of cell culture, cell viability was greater than 80%. This represents a substantial improvement over nanoparticles made solely of chitosan of high or low Mw that exhibited IC50 values of 36.4 and 18 µg of nanoparticles/

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cm2, respectively, when assayed using the same cell line (Köping-Höggard et al., 2008). This reduction in cytotoxicity has been attributed to the presence of hyaluronic acid, due to its high biocompatibility and involvement in biological processes such as cell adhesion and proliferation, among others. Cell transfection efficiency as assayed by image analysis of CLSM micrographs revealed that chitosan–hyaluronic acid nanoparticles had the ability to elicit high and long lasting transfection levels. The levels of gene expression up to four days showed significant differences with chitosan Mw, while beyond this point and up to ten days a plateau was achieved and no significant differences were observed. The effect of the chitosan of low Mw has been attributed to onset release of the pDNA from the nanoparticles. Such an expression profile has also been described for simple chitosan oligomers– pDNA polylelectrolyte complexes and attributed to the rapid dissociation of the pDNA (Köping-Höggard et al., 2004). High Mw chitosan shows greater affinity for DNA and the complexes are less easily dissociated than those of chitosan oligomers as a consequence of the cooperative nature of the binding between both polylectrolytes (Köping-Höggard et al., 2003). Yet another conclusion from the transfection studies on chitosan–hyaluronic acid nanoparticles was that the overall number of cells expressing a green fluorescent protein increased with the amount of hyaluronic acid included in the nanoparticles. The mechanisms whereby hyaluronic acid enhances gene transfection are yet to be fully elucidated, but it has been speculated that the interaction of hyaluronic acid with cellular-specific and non-specific mechanisms, its capacity to enter the cell nucleus and its functional role in cell signalling, may altogether account for its effect to promote gene transfection (De la Fuente et al., 2008b). Confocal laser scanning microscopy studies of the internalization and intracellular distribution conducted with fluorescentlylabeled pDNA (pβ-Gal), revealed that cells exposed to the nanoparticles showed an intense green fluorescence, corresponding to the labelled pDNA distributed in the cell cytoplasm. No differences among the various compositions were observed, thus suggesting that cellular uptake is not the limiting factor for efficient gene transfection. However, the visualization of the intracellular fate of the plasmid at 15 h after incubation of the pDNAloaded nanoparticles with the cells for 1 h, showed green spots associated with the labelled pDNA in the cytoplasm and also in the nuclei. Overall, the images of the patterns of intracellular distribution led us to conclude that the intracellular trafficking and, thus, the access of the pDNA to the nucleus are dependent on the nanoparticle composition. This interpretation would also be in agreement with the improved gene transfection capacity observed for nanoparticles made of polysaccharide oligomers. Recent proof-of-concept studies have also shown that this nanoparticle system is effective in transfecting the cornea and the conjunctiva after topical administration in rabbits (De la Fuente et al., 2008c).

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25.5.4 Chitosan konjac–glucomannan nanoparticles Methodologies and critical production parameters The study of the incorporation of konjac glucomannan and its phosphorylated derivative in chitosan-based nanoparticle systems was pioneered by our group under two distinct approaches: (a) Mixing an aqueous phase containing chitosan with an aqueous phase that contains the cross-linking agent TPP (i.e. formation of primary chitosan–TPP nanoparticles) followed by their coating with phosphorylated konjac glucomannan (Cuña et al., 2006). (b) Mixing an aqueous phase containing chitosan with an aqueous phase that contains konjac glucomannan with or without the crosslinking agent, TPP (Alonso-Sande et al., 2006a). Each of these approaches is briefly described below. Under approach (a), chitosan nanoparticles were obtained by ionic gelation of chitosan with TPP (chitosan/TPP ratios of 3/1 and 6/1). The incorporation of phosphorylated konjac glucomannan onto the nanoparticle structure was achieved by dissolving it into the nanoparticle suspending medium (Cuña et al., 2006). The heterotypic association in this co-gelled system is attributed to intermolecular hydrogen bonds (Xiao et al., 2000) along with electrostatic interactions between oppositely charged groups in phosphorylated konjac glucomannan derivative and free – NH 3+ groups in chitosan. Two different strategies were tested for coating chitosan–TPP nanoparticles with phosphorylated konjac glucomannan: in the first protocol, the chitosan nanoparticles were isolated before their coating; while for the second one, chitosan nanoparticles were not isolated, but coated with glucomannan in the presence of free chitosan. The size of nanoparticles obtained under both protocols ranged from approximately 170 to 300 nm and their zeta potential values were inverted from positive to negative as a consequence of the glucomannan coating. An advantage featured by nanoparticles harnessed under the two tested glucomannan coating strategies is that they could be freeze-dried in the presence or absence of cryoprotective agents, preserving their original characteristics. Moreover, studies addressing the stability of these nanoparticles in buffered media evidenced that the glucomannan coating played a key role in preventing the aggregation of the system. In accordance with approach (b), chitosan–konjac glucomannan and chitosan–phosphorylated konjac glucomannan nanoparticles were formed either in the absence or in the presence of added TPP; in the former case, it was possible to prepare nanoparticles when the chitosan/konjac glucomannan theoretical mass ratios were in the range 6/1 to 6/24 (Alonso-Sande et al., 2006a). The mechanism of formation is due to intermolecular hydrogen bonding between –OH groups of konjac glucomannan, and chitosan – NH 3+

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groups. It was found that incorporation of konjac glucomannan reaches a maximum at a 6/13.8 theoretical chitosan/konjac glucomannan ratio, beyond which glucomannan is no longer integrated into the nanoparticle structure. The increase in konjac glucomannan contents in the nanoparticles is accompanied by a concomitant decrease in hydrodynamic diameter that varies from ~800 to ~400 nm as the chitosan/konjac glucomannan ratio varies from 6/6 to 6/24. The zeta potential of these nanoparticles was close to neutral, irrespective of the chitosan/konjac glucomannan ratio, within the range of – 0.4 to –2.3 mV for the same range of composition. This has been explained as an effect of the plant polysaccharide that shields the charges on the chitosan. This system was also found to be under kinetic control during its formation and structural organization, by contrast with chitosan–phosphorylated konjac glucomannan nanoparticles that form spontaneously. In turn, chitosan– phosphorylated konjac glucomannan nanoparticles have a size range from ~180 to 310 nm depending in polymer composition (chitosan/phosphorylated konjac glucomannan mass ratio): the increase in phosphorylated konjac glucomannan ratio leads to a reduction in size. This effect is explained as a consequence of a condensation effect due to neutralization of amino charges by anionic phosphorylated konjac glucomannan, similar to that observed in chitosan–TPP nanoparticles (Janes et al., 2003). Despite the general effect of phosphorylated konjac glucomannan in reducing the nanoparticle size, positive zeta potential values (> +30 mV) persisted irrespective of the amount of the glucomannan component in the system, an effect attributed to the proportionally much greater charge density of chitosan (~85%) than that of phosphorylated konjac glucomannan (~7%). By contrast with the nanoparticles just described, in chitosan–TPP–konjac glucomannan and chitosan–TPP–phosphorylated konjac glucomannan hybrid nanoparticles, TPP is involved in the particle formation by ionic cross-linking. This does not lead to an overall change in their morphology, though a substantial difference in these nanoparticles is that the incorporation of phosphorylated konjac glucomannan did not cause a significant modification in particle size in the range of chitosan to phosphorylated konjac glucomannan ratio 6/1.2 to 6/3, which can be attributed to the low amount of incorporated glucomannan. The zeta potential is drastically reduced from +38.09 to +15.2 mV, as a consequence of the neutralization of the free positive amino groups remaining in the CS/TPP nanoparticles. A very important characteristic featured by chitosan–konjac glucomannan nanoparticles is that the stability in high ionic strength media such as those found in physiological conditions, is substantially improved with respect to chitosan–TPP nanoparticles, as is illustrated in Figure 25.7 for the evolution of particle size in PBS buffer pH 7.4 for various chitosan–TPP and chitosan– konjac glucomannan nanoparticle systems. This stabilization effect is conceived to be dominated by hydrophobic and hydrogen bonding forces.

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3500 3000

Size (nm)

2500 2000 1500 1000 500 0 0

30

60 Time (min)

90

120

25.7 Evolution of the nanoparticles’ size following their incubation in PBS pH 7.4 up to 2 h at room temperature (mean ± SD, n = 4). Nanoparticles polymer composition: Chitosan/phosphorylated konjac glucomannan (6/4.6) (filled triangles); chitosan/phosphorylated konjac glucomannan (6/13.8) (diamonds); chitosan/TPP/phosphorylated konjac glucomannan (6/1/4.6) (filled boxes); chitosan/konjac glucomannan (6/6) (empty triangles); chitosan/TPP/phosphorylated konjac glucomannan (6/1/1.8) (empty boxes). Source: Alonso-Sande et al. (2006). With permission of American Chemical Society.

Drug delivery and biopharmaceutical applications Chitosan–konjac glucomannan nanoparticles were developed with a view to improve the efficacy of chitosan nanoparticles for enhancing the intestinal absorption of peptides and proteins, such as insulin (Alonso-Sande et al., 2008b). The efficacy of these formulations to enhance the absorption of insulin was assessed by measuring the plasma glucose levels after administration to rats. The presence of konjac glucomannan in the nanostructure was critical in order to obtain a 50% decrease in glucose serum levels. Interestingly, chitosan–glucomannan nanoparticles were able to elicit a delayed hypoglycaemic response at 14 h post-administration, and this response was maintained for = 10h. The success of chitosan–glucomannan nanoparticles as compared with those made of chitosan could be related to the observed stabilising effect of glucomannan. The role of konjac glucomannan in the enhancement of the interaction of the nanoparticles with the intestinal epithelium has been further studied by in vitro tests using a co-culture of Caco-2 and Raji cells as a model to M-cells. This showed that the presence of konjac glucomannan enhanced drastically the uptake of nanoparticles to M-cells with respect to that of chitosan–TPP nanoparticles (Alonso-Sande et al., 2006b).

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Future trends

Throughout this chapter, we have intended to offer an overview on fundamental and applied aspects underlying the physicochemical properties of chitosanbased nanoparticle systems as well as their biopharmaceutical applications and drug delivery properties as vehicles of bioactive macromolecules. Specific emphasis has been given to nanoparticle systems formed by ionotropic gelation of chitosan alone and in hybrid co-gel systems comprising alginate, hyaluronic acid and konjac glucomannan of potential use in the biomedical field. Undoubtedly, chitosan nanoparticles have demonstrated an interesting potential for the administration of macromolecules and vaccines by mucosal routes. Currently ongoing toxicological and mechanistic studies will be the basis for their clinical application to be realized in the near future. Hybrid nanoparticles comprising chitosan and other polysaccharides will continue to widen the range of strategies for targeting specific mucosa tissues, as well as to reduce the necessary amount of chitosan to achieve the desired effect of the nanocarrier system. Novel nanometric advanced delivery systems based on co-gel chitosan with other polysaccharides, will undoubtedly also see developments related with layer-by-layer nanocoated structures, stimuli-responsive (‘smart’) delivery and surface-modified systems obtained by conjugation of non-synthetic moieties (sugars, enzymes, antibodies, lectins, folic acid) aimed for specific target delivery in cancer therapy by intravenous administration. To this end, systems fully biocompatible with blood will need to be developed.

25.7

Sources of further information and advice

The following books deal with nanotechnology and polymers in drug delivery: Nanoparticles for Pharmaceutical Application (Domb et al., 2007). Nanoparticle Drug Delivery Systems (Thassu et al., 2007). Nanotheraputics: Drug Delivery Concepts in Nanoscience (Lamprecht, 2007). Nanotechnologies for Cancer Therapy (Amiji, 2007). Nanoparticles as Drug Carriers (Torchilin, 2006). Polymers in Drug Delivery (Ucheqbu and Schatzlein, 2006). Biological and Pharmaceutical Nanomaterials (Nanotechnologies for the Life Sciences) (Kumar, 2006). Biomedical Nanotechnology (Malsch, 2005). Biomaterials for Delivery and Targeting of Proteins and Nucleic Acids (Mahato, 2005). Carrier-Based Drug Delivery (Sonke, 2004). New Trends in Polymers for Oral and Parenteral Administration: From Design to Receptors (Barratt et al., 2001).

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Trade/professional bodies of interest American Chemical Society www.acs.org Royal Society of Chemistry www.rsc.org Controlled Release Society www.controlledreleasesociety.org Association of Pharmacie Galénique Industrielle www.apgi.org European Federation for Pharmaceutical Sciences http://www.eufeps.org/ Sociedad Hispano Lusa de Liberación de Fármacos www.splc-crs.org/es/ index.htm Red de Sistemas de Liberación de Moléculas Activas www.redslma.com European Chitin Society www.euchis.org Iberoamerican Chitin Society www.siaq.net ASME Nanotechnology Institute http://www.nanotechnologyinstitute.org/ index.shtml International Association of Nanotechnology http://www.aanano.com/ index.html Australian Research Council Nanotechnology Network http://www.ausnano.net/network/TNN/32/

Institutions The Institute of Nanotechnology http://www.nano.org.uk/aboutus/activities.htm Foresight Nanotech Institute http://www.foresight.org/ Center for Biological and Environmental Nanotechnology (Rice University) http://cben.rice.edu/ Center for Nanomaterials Research (Dartmouth College) http://engineering.dartmouth.edu/nanomaterials/ Center for Nanotechnology (University of Washington) http://www.nano.washington.edu/index.asp CNT Center for Nanotechnology http://www.ipt.arc.nasa.gov/ NanoScience and Technology Center http://nsti.org/ Nanobiotechnology Center (Cornell University) http://www.nbtc.cornell.edu/ Center for Molecular Nanofabrication and Devices (Penn State University) http://www.cmnd.psu.edu/ Center of Excellence for Nanotechnology and Biotechnology http://www.metucenter.metu.edu.tr/index.htm Center for Nano and Molecular Science and Technology (University of Texas at Austin) http://www.cnm.utexas.edu/nstinformation.html Nanoscale Science and Engineering Center (Columbia University) http://www.cise.columbia.edu/nsec/ Science of Nanoscale Systems and their Device Applications (Harvard University) http://www.nsec.harvard.edu/ Nanoscale Science and Engineering Center for Integrated Nanopatterning and Detection Technologies (Northwestern University) http:// www.nsec.northwestern.edu/ © 2008, Woodhead Publishing Limited

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Nanoscale Science and Engineering Center for Directed Assembly of Nanostructures (Rensselaer Polytechnic Institute) http:// www.nsec.northwestern.edu/ Center for Nanoscience (Ludwig Maximillians University) http://www.cens.de/ Center for Functional Nanostructures (Karlsruhe University) http:// www.cfn.uni–karlsruhe.de/about.html Center for Nanotechnology (Muenster) http://www.centech.de National Nanotechnology Initiative http://www.nano.gov/ NIST Center for Nanoscale Science and Nanotechnology http://cnst.nist.gov/ Center for Functional Nanomaterials (Brookhaven National Laboratories) http://www.bnl.gov/cfn/ Microsystems and Nanotechnology Center (Cranfield University) http:// wwwlegacy.cranfield.ac.uk/sas/materials/nanotech/ Nanochemistry http://www.nanochem.kth.se/nano/ Nanotechnology Research Institute (Japan) http://unit.aist.go.jp/nanotech/ index.html Competence Center for the Application of Nanostructures in Optoelectronics http://www.nanop.de/ Pacific Northwest National Laboratory http://www.pnl.gov/nano/

Research and interest groups European Project Nanobiosaccharides www.nanobiosaccharides.org Asia Pacific Nanotechnology Forum http://www.apnf.org/ Excellence Network NanoBiotechnology http://www.ennab.de/ The Nanotechnology Group http://www.thenanotechnologygroup.org/ Websites with updated news on nanotechnology and pharmaceutical technology www.in–pharmatechnologist.com NanotechWeb http://nanotechweb.org/ MEMS and Nanotechnology Clearinghouse http://www.memsnet.org/ Nanoforum – European Nanotechnology Gateway http://www.nanoforum.org/ Nanotechnology Now http://www.nanotech–now.com/ The International Nanotechnology Business Directory http:// www.nanovip.com/ The Nanotube Site http://www.pa.msu.edu/cmp/csc/nanotube.html Nanotechnology Today http://www.geocities.com/aardduck/ nanotech_today.html Nanotechnology.com http://www.nanotechnology.com/ NanoNed (Nanotechnology Network in the Netherlands) http://www.nanoned.nl/default.htm NanoNet http://www.nanonet.org.uk/ Nanomedicine http://www.nano–biology.net/ NanoTsunami http://www.nano–tsunami.com/ NanoBionet http://www.nanobionet.de/nanobiotechnology.htm © 2008, Woodhead Publishing Limited

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Nano2Life (European Network of Excellence in Nanobiotechnology) http://www.nano2life.org/ NanoBioLab (Universitat des Saarlandes) http://www.uni–saarland.de/fak8/hempelmann/nanobiolab.htm Nanopolis http://www.nanopolis.net/ Nanotechnology Database http://www.wtec.org/loyola/nano/links.htm Nanotechnology Researchers Network of Excellence (Japan) http:// www.nanonet.go.jp/english/ Nanotechnology Websites http://www.hyperion.ie/Nano.htm

25.8

Acknowledgements

Financial support of the European Union from the Nanobiosaccharides project (Ref No. 013882 of call FP6–2003–NMP–TI–3–Main) is gratefully acknowledged.

25.9

References

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Lamprecht A (2007), Nanotherapuetics: Drug Delivery Concepts in Nanoscience, Hackensack, NJ: World Scientific Publishing Co. Inc. Lee S T, Mi F L, Shen Y J and Shyu S S (2001), ‘Equilibrium and kinetic studies of copper(II) ion uptake by chitosan-tripolyphosphate chelating resin’, Polymer, 42, 1879– 1892. Lehr C M, Bouwstra J A, Schacht E H and Junginger H E (1992), ‘In-vitro evaluation of mucoadhesive properties of chitosan and some other natural polymers’, Int J Pharm, 78, 43–48. López-León T, Carvalho E L S, Seijo B, Ortega-Vinuesa J L and Bastos-González D (2005), ‘Physicochemical characterization of chitosan nanoparticles: electrokinetic and stability behavior’, J Colloid Interf Sci, 283, 344–351. Loretz B and Bernkop-Schnürch A (2006), ‘In vitro evaluation of chitosan–EDTA conjugate polyplexes as a nanoparticulate gene delivery system’, AAPS Journal, 8(4), art. no. 85. Ma Z, Yeoh H H and Lim L Y (2002), ‘Formulation pH modulates the interaction of insulin with chitosan nanoparticles’, J Pharm Sci, 91, 1396–1404. Ma Z, Lim T M and Lim L Y (2005), ‘Pharmacological activity of peroral chitosaninsulin nanoparticles in diabetic rats’, Int J Pharm, 293, 271–280. Maeda M, Shimahara H and Sugiyama N (1980), ‘Detailed examination of the branched structure of konjac glucomannan’, Agric Biol Chem, 44(2), 245–252. Maestrelli F, Garcia-Fuentes M, Mura P and Alonso M J (2006), ‘A new drug nanocarrier consisting of chitosan and hydoxypropylcyclodextrin’, Eur J Pharm Biopharm, 63, 79–86. Mahato R I (2005), Biomaterials for Delivery and Targeting of Proteins and Nucleic Acids, Boca Raton, F L: CRC Press. Malsch N H (2005), Biomedical Nanotechnology, Boca Raton, FL: CRC Press Taylor & Francis Group. Mao H Q, Roy K, Troung-Le V L, Janes K A, Lin K Y, Wang Y, August J T and Leong K W (2001), ‘Chitosan-DNA nanoparticles as gene carriers: synthesis, characterization and transfection efficiency’, J Control Rel, 70, 399–421. Marty J J, Oppenheim R C and Speiser P (1978), ‘Nanoparticles – a new colloidal drug delivery system’, Pharm Acta Helv, 53(1), 17–23. McClean S, Prosser E, Meehan E, O’Malley D, Clarke N, Ramtoola Z and Brayden D (1998), ‘Binding and uptake of biodegradable poly-DL-lactide micro- and nanoparticles in intestinal epithelia’, Eur J Pharm Sci, 6(2), 153–163. Mi F L, Shyu S S, Lee S T and Wong T B (1999a), ‘Kinetic study of chitosan-tripolyphosphate complex reaction and acid–resistive properties of the chitosan-tripolyphosphate gel beads prepared by in-liquid curing method’, J Polym Sci Pol Phys, 37, 1551– 1564. Mi F L, Shyu S S, Lee S T, Kuan C Y, Lee S T, Lu K T and Jang S F (1999b), ‘Chitosan– polyelectrolyte complexation for the preparation of gel beads and controlled release of anticancer drug. I. Effect of phosphorous polyelectrolyte complex and enzymatic hydrolysis of polymer’, J Appl Polym Sci, 74, 1868–1879. Mi F L, Shyu S S, Wong T B, Jang S F, Lee S T and Lu K T (1999c), ‘Chitosan– polyelectrolyte complexation for the preparation of gel beads and controlled release of anticancer drug. II. Effect of pH–dependent ionic crosslinking or interpolymer complex using tripolyphosphate or polyphosphate as reagent’, J Appl Polym Sci, 74, 1093– 1107. Milas M and Rinaudo M (2004), ‘Characterization and properties of hyaluronic acid

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Vandevord P J, Matthew H W T, Desilva S P, Mayton L, Wu B and Wooley P H (2002), ‘Evaluation of the biocompatibility of a chitosan scaffold in mice’, J Biomed Mat Res, 59(3), 585–590. Vårum K M, Anthonsen M W, Grasdalen H and Smidsrød O (1991), ‘Determination of the degree of N-acetylation and the distribution of N-acetyl groups in partially Ndeacetylated chitins (chitosans) by highfield n.m.r. spectroscopy’, Carbohyd Res, 211, 17–23. Vila A, Sánchez A, Tobío M, Calvo P and Alonso M J (2002), ‘Design of biodegradable particles for protein delivery’, J Control Rel, 78, 15–24. Vila A, Sánchez A, Janes K A, Behrens I, Kissel T, Vila-Jato J L and Alonso M J (2004), ‘Low molecular weight chitosan nanoparticles as new carriers for nasal vaccine delivery in mice’, Eur J Pharm Biopharm, 57, 123–131. Vinogradov S V, Bronich T K and Kabanov A V (2002), ‘Nanosized cationic hydrogels for drug delivery: preparation, properties and interactions with cells’, Adv Drug Del Rev, 54, 135–147. Wu Y, Yang W, Wang C, Hu J and Fu S (2005), ‘Chitosan nanoparticles as a novel delivery system for ammonium glycyrrhizinate’, Int J Pharm, 295, 235–245. Xiao C, Gao S, Wang H and Zhang L (2000), ‘Blend films from chitosan and konjac glucomannan solutions’, J Appl Polym Sci, 76(4), 509–515. Xu S, Yamanaka J, Sato S, Miyama I and Yonese M (2000), ‘Characteristics of complexes composed of sodium hyaluronate and bovine serum albumin’, Chem Pharm Bull, 48(6), 779–783. Xu Y and Du Y (2003), ‘Effect of molecular structure of chitosan on protein delivery properties of chitosan nanoparticles’, Int J Pharm, 250, 215–226. Zhang H, Oh M, Allen C and Kumacheva E (2004), ‘Monodisperse chitosan nanoparticles for mucosal drug delivery’, Biomacromolecules, 5, 2461–2468.

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Part VI Biocompatibility of natural-based polymers

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26 In vivo tissue responses to natural-origin biomaterials T. C. S A N T O S, A. P. M A R Q U E S and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

26.1

Introduction

Increasing life expectancy allied with life welfare has been contributing to the progress of biotechnology. In fact the development of biomaterials answers to the rising needs for new tissue replacement/regeneration strategies. Nonetheless, the implantation of biomaterials leads to the development of immune responses that may go from a light inflammatory reaction to severe tissue damage and ultimately rejection of the implant. Following inflammation and immune reactions, a variety of mediators are released, inducing the recruitment of subpopulations of cells that, if not properly regulated, can cause tissue damage. Those mediators are released by blood platelets and by several types of cells, such as tissue mast cells, leukocytes, fibroblasts, endothelial cells, osteoblasts and osteoclasts. Nevertheless, despite all the damage that severe and chronic inflammatory reactions may cause to the host and implanted material, the initial acute inflammation is also essential for the initiation of healing and regeneration of new tissue. One of the biggest challenges in the development of implantable biomaterials is the manipulation of these materials to enhance their in vivo performance minimizing host reactions. In fact, the type of host tissue reaction that develops following biomaterials implantation depends on the surface physical and chemical characteristics of the material and device, but also on the type of tissue involved and its mechanical function, and on the general host physiological condition. This chapter aims to offer an overview of the inflammatory and immune processes and their contextualization within host responses triggered by the implantation of natural-origin biomaterials. Moreover, the challenges faced in the assessment of tissue responses to natural-origin biomaterials will be focused upon. The animal models currently used to test tissue responses, as well as the available and future strategies to control or enhance those responses, depending on the aim and function of the developed natural-origin biomaterial will be reviewed. 683 © 2008, Woodhead Publishing Limited

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Inflammation and foreign-body reactions to biomaterials

After implantation of a medical device, the host tissue will inevitably be traumatized by the implantation procedure (Mikos et al., 1998; Hunt, 2001; Stevens et al., 2002; Williams, 2001) triggering an inflammatory response. The consequent recruitment of cells mimics the one observed in a local inflammation (Spargo et al., 1994), and the analysis of this inflammatory response, together with the response to trauma is, therefore, considered critical for the overall biocompatibility assessment (Hunt, 2001). Although the implantation of a foreign-body elicits a host response towards the implant with the features of a chronic inflammation, there is always an early acute inflammatory response, mainly endorsed to the implantation procedure (Figure 26.1). The assembling of an acute inflammatory response may take place in minutes or hours, depending on the severity of the injury and usually lasts hours or days (Fantone and Ward, 1999; Goldsby et al., 2000). Essentially, the purposes of inflammation are to destroy (or contain) the damaging agent,

Implantation procedure

Acute inflammation • Cells • Molecules

Tissue damage

Chronic inflammation

Surface adsorption • Fibrin • Complement proteins • Antibodies

Onset of foreignbody reaction

26.1 Scheme simplifying the events triggered by the implantation of a biomaterial, until the development of a chronic inflammation, which, eventually, may result in rejection of the implant.

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to initiate the repair process, and to return the damage tissue to useful function as a continuous event (Fantone and Ward, 1999; Stevens et al., 2002). As a wound is created, coagulation will take place in the context of acute inflammation. Simultaneously to the activation of the coagulation cascade, the complement system, which has the capability of distinguishing ‘self’ from ‘non-self’ (Atkinson and Farries, 1987; Mollnes, 1997) is activated (Gorbet and Sefton, 2004; Stevens et al., 2002). In biomaterials implantation, the complement system may be activated either by the classical pathway, through the interaction of plasma proteins, such as immunoglobulins (Williams, 2001) and fibrin (Williams, 2001; Stevens et al., 2002; Gorbet and Sefton, 2004, Nilsson et al., 2007), with the surface of the material, or by the alternative pathway, with the inadequate down-regulation of convertase (Nilsson et al., 2007). Besides complement system activation, the adsorbed proteins onto the surface of the implanted materials act as a strong chemoatractant to polymorphonuclear neutrophils (PMNs). After being recruited to the site of injury and/or implantation, PMNs become activated, undergo a ‘respiratory burst’, which generates reactive oxygen species, and degenerate (Williams, 2001; Stevens et al., 2002). Therefore, neutrophils are the dominant cell type in the early phase of acute inflammation. Nonetheless, within 24 hours blood monocytes, under the influence of chemotactic factors, begin to migrate into the damaged tissue and after 48–72 hours they are the predominant cell type (Spargo et al., 1994; Stevens et al., 2002). Macrophages derived from blood monocytes continue the phagocytic work initiated by neutrophils (Bellingan et al., 1996; Stevens et al., 2002), although they might also act as antigenpresenting cells (APCs), after processing the material (Williams, 2001), instigating specific immunological responses (Stevens et al., 2002) in which lymphocytes also participate (Stevens et al., 2002). The outcome of the inflammation process depends on a variety of factors, including the nature and destructibility of the injurious agent, the extent of tissue damage and the properties of the tissue in which the damage has occurred (Stevens et al., 2002). Furthermore, the presence of a biomaterial constitutes an additional factor to disrupt the normal course of the inflammation and healing processes. In general, it is accepted that the hallmark of acute inflammation is the interaction of recruited leukocytes with the proteins adsorbed to the surface of the material, and consequent reaction to the implanted biomaterial, while the formation of foreign-body giant cells (FBGCs) usually indicates the transition to a chronic inflammatory process (Hunt, 2001, Anderson, 2000). Nonetheless, the same features may co-exist, attesting the development of acute and chronic inflammation (Figure 26.1), simultaneously (Lickorish et al., 2004). The prolonged presence of the implanted biomaterial, as well as its degradability are, per se, strong features regarding the formation of an abscess as a component of chronic inflammatory reaction (Griffiths et al., 1996;

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Hu et al., 2001). By definition, this chronic inflammation lasts for weeks or months and its brand characteristics are ongoing tissue damage, often caused by the inflammatory cells in the infiltrate, a chronic infiltrate and fibrosis (Goldsby et al., 2000). The presence of non-degradable foreign materials in the tissues may also stimulate a chronic granulomatous inflammatory response (Goldsby et al., 2000; Luttikhuizen et al., 2006; Stevens et al., 2002; Junqueira and Carneiro, 2005; Lickorish et al., 2004; Williams, 2001). Some implanted foreign materials are refractile when viewed with polarized light and thus, can be easily identified within the granulomas or giant cells (Liao et al., 2000). Nevertheless, chronic inflammation may also occur immediately after the implantation of some biomaterials. The mechanisms of this non-immune type of chronic inflammation are not really clear, but some foreign materials may activate macrophages to release mediators that induce an early inflammatory reaction with fibrosis (Luttikhuizen et al., 2006).

26.3

Role of host tissues in biomaterials implantation

The nature/type of the tissue where the implant is allocated plays an important role either in the initiation of the inflammatory process (acute inflammation), or in the progression into a severe inflammatory reaction. In the biomaterials field it is assumed that any type of tissue from the human body may once be recipient of an implant. Therefore, it is mandatory to consider the type of cells, the extracellular matrix and the whole network of tissues involved, as the incoming environment for the foreign implant, as well as the protein content of the tissue, which is susceptible to adhere to the surface of the implant. Actually, independently of the final function of a biomaterial/device, it will always have very close contact with whole blood. Therefore, the blood biocompatibility of any developed biomaterial constitutes a critical issue to assess. Moreover, the resolution or repair of injury also depends on the type of tissue where the biomaterial is implanted. The proliferative capacity of the tissue, the extent of injury and the persistence of the tissue framework at the implant site are crucial for controlling either: (a) the regeneration of tissue-specific parenchymal cells and restitution of the normal tissue structure or; (b) the reorganization and replacement of the injured tissue with newly synthesized fibrovascular connective tissue (Mikos et al., 1998). In this context, this section will give special attention to the histological features of connective tissue, muscle and blood, due to its role in the animal models that will be discussed.

26.3.1 Connective tissue: The origin Connective tissue is considered the basis of all tissues, since it is present in all of them, providing structural support and maintaining the form of the

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whole body. Despite this, the different forms of connective tissue differ in the type of cells and in the produced extracellular matrix (Junqueira and Carneiro, 2005; Stevens et al., 2002). All cells constituting the connective tissue originate from mesenchyme cells, which are derived, in their majority, from the middle germ layer of the embryo, the mesoderm. The most widespread cell type of the connective tissue is the fibroblast (Stevens et al., 2002). Macrophages are present in the connective tissue with the function of phagocytose microorganisms, parasites, foreign bodies and cell debris. They are derived from blood monocytes which migrate into connective tissue, thus differentiating into tissue macrophages. In the different tissues this cell type is given different names, such as alveolar macrophages in the lung, Kupffer cells in the liver and osteoclasts in the bone. In the presence of a foreign-body, several macrophages may fuse together originating a multinuclear foreign-body giant cell (FBGC). These cells are often seen at the onset of inflammatory responses (Wheater et al., 1979), at the sites of implantation of biomaterials, as mentioned in the previous section. Other cells present in the connective tissue and also important in the inflammatory response are mast cells. These cells show an oval or round shape and the cytoplasm is filled with large basophilic granules, which are released in inflammatory responses, namely in hypersensitivity reactions (Metz and Maurer, 2007), and can be found predominantly adjacent to blood vessels. The antibody production in connective tissue is the responsibility of plasma cells. Plasma cells derive from B-lymphocytes and release the antibodies that bind to the antigens in the course of immune responses. These cells are large, with eccentric nuclei, basophilic cytoplasm containing abundant rough endoplasmic reticulum (RER) and well-developed Golgi bodies, and can be found in sites of chronic inflammation or in sites of high risk of bacterial or foreign protein invasion (Wheater et al., 1979; Goldsby et al., 2000). The extracellular matrix of the connective tissue is constituted by: (a) protein fibres (collagen fibres, reticular fibres and elastic fibres) differing, namely, in the amino acids content and thickness; (b) ground substance, which is an amorphous, transparent material composed mainly of water, glycoproteins and proteoglycans; and (c) tissue fluid (Goldsby et al., 2000). All these components are potential adsorbents to the surface of implanted foreign materials, thus leading to the development of the host response after cell interaction.

26.3.2 Muscle In mammals there are three types of muscular tissue, differing in morphological and functional characteristics: skeletal muscle, with quick and voluntary contraction, is composed of long cylindrical multinucleated cells with crossstriations (the muscle fibres) (Heffner Jr, 1992; Hays and Armbrustmacher,

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1999); cardiac muscle, constituted by elongated, branched individual cells that lie parallel to each other, has also cross-striations, but the contraction is involuntary, vigorous and rhythmic (Billingham, 1992); and smooth muscle composed by fusiform cells without cross-striations and promoting a slow and involuntary contraction (Junqueira and Carneiro, 2005). Concerning the evaluation of the host response to biomaterials and tissue engineering constructs, probably the most relevant type of muscle tissue is the skeletal muscle or striated muscle. In fact, one of the most used strategies to assess the in vivo biological response of newly developed biomaterials is the intramuscular implantation of the materials in different animals (Liao et al., 2000; Meinel et al., 2005; Ravin et al., 2001). The skeletal muscle comprises a wide network of cells forming various structures, which have to be fed with nutrients through an extensive vascular network. For this reason, the muscle becomes a useful tissue to study the influx of circulating inflammatory cells and molecules to the site of implantation.

26.3.3 Skin Skin is the largest, the heaviest and the main cover organ of the body. It is constituted mainly by the epidermis (of ectodermal origin in the embryo) and by the dermis (of mesodermal origin in the embryo). Hypodermis, or subcutaneous tissue, comprising a subcutaneous layer of connective tissue and adipose tissue (panniculus adiposus), is located beneath the dermis (Harrist et al., 1999). The epidermis is constituted by a stratified squamous keratinized epithelium, containing keratinocytes, and by the melanocytes, Langerhans cells and Merkel’s cells. It consists of five layers of keratinocytes: from the dermis outward, stratum basale (stratum germinativum), stratum spinosum, stratum granulosum, stratum lucidum and stratum corneum. Melanocytes, specialized cells in producing melanin, are round cells with long and irregular extensions forming invaginations. Their cytoplasm typically contains several small mitochondria, well-developed Golgi complex and short RER. Langerhans cells are star-shaped bone marrow-derived cells which are able to bind, process and present antigens to T lymphocytes. Merkel’s cells are present, generally, in the thick skin of palms and soles, have small granules in the cytoplasm and serve as mechanoreceptors (Urmacher, 1992). Commonly, dermis consists of the connective tissue between the epidermis and the subcutaneous tissue (hypodermis). Its thickness varies according to the region of the body and possesses many projections (dermal papillae), which increase and reinforce the dermal-epidermal junction. The two layers of the dermis are the papillary layer, composed of loose connective tissue, fibroblasts, mast cells and macrophages, and the reticular layer, constituted by irregular dense connective tissue with more fibres and fewer cells, which confer elasticity to skin. The hair follicles, the sweat and the sebaceous

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glands, as well as the majority of skin nerves, are also present in the dermis (Urmacher, 1992; Harrist et al., 1999). Besides Langerhans cells present in the epidermis, cellular components of dermis such as resident lymphocytes, mast cells and macrophages have a critical role in defending the organism against invaders.

26.3.4 Blood As an incision is made for the implantation of any biomaterial, a rupture of blood vessels occurs and the contact between blood and the foreign material is unavoidable. Blood is composed of cells (erythrocytes and leukocytes) and platelets, and of plasma, an aqueous solution containing inorganic salts, hormones, vitamins, amino-acids, lipoproteins and proteins, including albumins, gamma globulins and fibrinogen, where the cells are suspended. Platelets derive from the cytoplasm of megakaryocytes of bone marrow. They are small with a core of small granules and their main role involves thrombus formation by aggregating in the site of injury (Johnson et al., 1999). After blood clotting, plasma deprived from fibrinogen and other clotting agents forms the blood serum (Goldsby et al., 2000). Leukocytes are the cellular blood component with key roles in the host responses to foreign bodies. The leukocytes are divided into granulocytes (neutrophils, basophils and eosinophils), fully differentiated cells characterized by the irregular segmented nucleus and the specific cytoplasmic granules, and lymphocytes and monocytes, characterized by their regular nucleus and non-specific cytoplasmic granules. Neutrophils (PMNs), the most abundant leukocytes in circulation, are the first line of defence against foreign invaders due to their phagocytic ability. Additionally, they play a critical role at injury sites or wounds by releasing the components of their specific granules such as phagocytins, lysosomal enzymes and peroxidase. Eosinophils also contribute to the cocktail of enzymes released at implantation sites, delivering substances such as acid phosphatase, cathepsin, among others. These cells are also involved in selective phagocytosis and are highly active in allergic reactions. Basophils, in turn, although with a high amount of histamine and heparincontaining granules in their cytoplasm, specifically act in response to antigens (Goldsby et al., 2000). Monocytes, although not granulocytes, have the ability to differentiate into phagocytic cells, the macrophages (Junqueira and Carneiro, 2005), and to release degradative molecules (Khouw et al., 2000c), thus being involved in similar defence reactions. Lymphocytes are the hallmark of immune responses. These small cells with dense and regular nucleus and small cytoplasm are classified as B, T or natural killer (NK) cells. The B and T cells are the only cells capable of selectively recognizing a specific epitope among a vast number. These two

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types of cells differ in their life history, surface receptors and behaviour during an immune response. The origin of all lymphocytes is the bone marrow, where B lymphocytes and NK cells mature and become functional. After leaving the bone marrow, these cells enter blood circulation and colonize connective tissues, epithelia, lymphoid nodules and lymphoid organs. Within tissues, B lymphocytes are able to recognize an epitope, to proliferate and to redifferentiate into plasma cells, which secrete antibodies against that recognized epitope. In contrast, T lymphocytes mature in the thymus, where the T lymphocyte precursors arrive from bone marrow, and are ‘distributed’ (as CD4+ or T helper lymphocytes, and CD8+ or cytotoxic T lymphocytes) throughout the body connective tissues and lymphoid organs. B and T cells have the capacity to migrate from the tissues to the blood circulation and vice-versa. NK cells lack the surface markers that characterize B and T lymphocytes; therefore these cells do not need previous stimulation to attack virus-infected cells, transplanted cells and cancer cells (Johnson et al., 1999; Goldsby et al., 2000).

26.4

Assessing the in vivo tissue responses to natural-origin biomaterials

Assessment of the biological response to a newly developed biomaterial or biomedical device is not a recent concern in the biomaterials field. However, a better characterization of such response, at the cellular and molecular level, had been more extensively investigated in the last decades (Griffiths et al., 1996; Hunt et al., 1997; Hunt and Williams, 1995; Kao and Lee, 2001). Moreover, the complexity of the in vivo responses to implanted biomaterials renders this assessment a challenging issue to address. Different methodologies can be used to identify the recruited inflammatory cells, as well as to elucidate their enrolment time schedule (Spargo et al., 1994). Nevertheless, the most currently used animal models are the subcutaneous (Khouw et al., 2000a; Brodbeck et al., 2003; Marques et al., 2005; Lickorish et al., 2004), the intramuscular (Ravin et al., 2001; Meinel et al., 2005) and the intraperitoneal (Usami et al., 1998; Robitaille et al., 2005) implantations. All these models, except perhaps subcutaneous implantation, provide a good exposition of biomaterials to whole blood and, therefore are also suitable for testing the blood biocompatibility. Besides blood cellular components, there are several blood-related factors such as the complement system (Mollnes, 1998) and the coagulation cascade proteins (Hunt, 2001), as well as the adhesion proteins adsorbed to the surface of the biomaterial upon implantation that influence the host reaction (Tang et al., 1996; Nimeri et al., 2002). These factors dictate, to some extent the subsequent cell adhesion, differentiation and proliferation (Szaba and Smiley, 2002) and host response. Nonetheless, and according to a recent publication (Ratner,

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2007), accurate assessment of the blood compatibility of a biomaterial is still not possible since it is a non-understandable issue that lacks standardization. Transgenic animals, knockout or knockin, generated by introducing a stable, in vitro recombined, foreign DNA sequence into their germline, constitute a type of animal model very useful to understand the role of some inflammatory/anti-inflammatory molecules and adhesion molecules involved in biomaterials implantation, since the engineered genetic modifications can generally originate a ‘gain’ or ‘loss’ of function (Ruelicke et al., 2007). The intraperitoneal cavity is an ‘open’ cavity, theoretically without air and in contact with the outer layers of the intraperitoneal organs, such as bowel, stomach, liver, spleen, among others. This cavity allows the recruitment of leukocytes from different sources, not only from the circulating blood but also from neighbouring organs that may react to the implantation of the foreign biomaterial (Robitaille et al., 2005). In what concerns the intramuscular model, as the skeletal muscle is a very well vascularized tissue that provides an easy and fast affluence of circulating inflammatory cells to the site of implantation, it constitutes a useful model to study the kinetics of inflammatory cell recruitment, its activation (Liao et al., 2000) and in particular the production of specific inflammatory/anti-inflammatory cytokines (Meinel et al., 2005), as well as to clarify the response of the tissue cells themselves (Meinel et al., 2005). In a subcutaneous implantation, the material is not only in contact with the components of the deeper layers of the skin mentioned above, but also with a portion of smooth muscle. In terms of surgical procedure, this is a very simple and easy model to test the in vivo reaction to biomaterials (Lickorish et al., 2004), providing useful information concerning the reaction of different cell types, such as the skin resident inflammatory cells (Spargo et al., 1994). Natural-based biomaterials are mainly constituted by proteins or polysaccharides which, under some circumstances, may be recognized either as natural invaders (such as bacteria), or as a body component. Furthermore, some authors (Ratner, 2007) consider that natural-based polymers may offer the key to producing biomaterials with better blood compatibility. Therefore, understanding the assembled host response to the very different naturalbased biomaterials always represents a big challenge and an immense effort for the immunology researchers in the field of biomaterials and tissue engineering. In the in vivo biocompatibility analysis of biomaterials, it is mandatory to have in mind at which level the evaluation has to be made. The main factors that influence the in vivo responses to natural-origin biomaterials are listed below: •

natural source of the biomaterial (type of natural polymer);

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size, shape and mechanical properties; physicochemical surface characteristics; degradation rate; animal model used for the reaction test; type of implantation procedure; general condition of the host.

In fact, the majority of the factors must not be considered independently. The shape and size of the biomaterial to be tested, as well as its final application, are important features to have in consideration when choosing the animal model. For example, for a compact or a scaffold material, subcutaneous or intramuscular implantation will be more suitable than intraperitoneal implantation. This type of model would be more appropriate to test the reaction of materials suspended in solutions, such as microparticles or nanoparticles. The final intended use and function of the implanted biomaterial is also related to the degradability issue. Generally, natural polymers undergo enzymatic degradation (Ali et al., 1994), and the degradation rate of a biomaterial is also linked to the type of response elicited by the host tissues. Phagocytic cells are normally able to remove debris from the tissue by engulfment and digestion, making the digestion of implanted materials an important issue to consider (Ali et al., 1994). In some cases it is not the biomaterial itself that induces a specific reaction, but the degradation products resulting from the concomitant action of the cells in the device. For example, hydrophobic surfaces potentially activate complement system, inducing higher degradation rates (Nilsson et al., 2007) although that activation may also be triggered by the plasma proteins that immediately bind to and cover the surface of a biomaterial after contact with blood or other body fluids (Nilsson et al., 2007). The tissue response elicited by an implanted biomaterial may also vary in different species. Khouw and co-workers (Khouw et al., 2000a), showed that the foreign-body reaction to subcutaneously implanted dermal sheep collagen differs between rats and mice, namely concerning the kinetics of inflammatory cell recruitment and phagocytosis. In this study, it was also shown that rats were able to mount a foreign-body reaction more effectively than mice. Contrarily, stroma formation and calcification were more abundant in mice compared with rats (Khouw et al., 2000a). High variance between animals in the same experiment is also a rather usual observation. Therefore, a statistically representative approach, not only in the number of implanted materials, but also in the number of attested animals, is crucial to reduce the standard deviation of the results of the experiment, and to have confidence in the tissue response of that particular species to the implanted biomaterials. A critical issue related to the immune reactions has to do with recurring contact with immunogenic items. In an interesting model of repetitive

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subcutaneous implantations of cross-linked collagens in rats, van Luyn and colleagues (van Luyn et al., 2001) showed that the animals became more reactive to the second challenge. The foreign-body reaction, namely the attraction of plasma cells, was enhanced after the second challenge with the natural-based biomaterials (van Luyn et al., 2001) which might be highly problematic, for example, when using the same type of biomaterial in different application sites in the same patient. The increased discussion regarding the number of animals used in research, led to the establishment of models that avoid animal sacrifice and limited data outcome. Ho and co-workers (Ho et al., 2007) were able to asses the real-time in vivo inflammatory response to a subcutaneous implant of genipi-cross-linked gelatine by in vivo bioluminescence in a transgenic mouse model carrying the luciferase gene driven by NF-κB-responsive elements. The movement of host molecules is, in fact, an important issue to consider in the monitoring of the inflammatory/ immune reaction to implanted biomaterials. In that particular case, the nuclear factor-κB (NF-κB) is a nuclear transcription factor, critically involved in the regulation of inflammatory cytokine production and, consequently, in inflammation (Bonizzi and Karin, 2004).

26.5

Controlling the in vivo tissue reactions to natural-origin biomaterials

The first approach in modulation of the host response to a newly developed biomaterial is modification at the level of the biomaterial, which mainly includes surface modification in terms of physicochemical characteristics. However, some modulation may also be performed at the host level. In terms of modulation at the materials level, coating of synthetic polymers with an external layer of natural polymers, such as chitosan or gelatin (Ciardelli and Chiono, 2006) is one of the followed strategies. Blending a synthetic polymer, polycaprolactone (PCL) with chitosan enhanced the system biocompatibility and biomimetics and hastened the degradation rate (Ciardelli and Chiono, 2006). The chemical groups present on the surface of the implanted material are also important for the mounted host response. The complement system is strongly activated by hydrophobic surfaces and by surfaces containing chemical groups such as NH2, OH or COOH, in comparison with hydrophilic surfaces (Nilsson et al., 2007). The strategy involving the modulation of the inflammatory/immune response at the host level usually concerns either the modulation of the host proteins adsorbed to the materials surface, or the control of appropriate host defence cells and of inflammatory/anti-inflammatory molecules. Furthermore, there are several host molecules with proved influence in the modulation of the inflammatory/immune responses, which may have potential to be used in the control of the responses to natural-origin biomaterials.

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In several studies, van Luyn and co-workers showed that depletion of macrophages (van Luyn et al., 1994) and deficiency in T-cells (van Luyn et al., 1998) inhibits the development of foreign-body reactions to cross-linked collagen in rats. Nonetheless, the in vitro pre-degradation and impregnation with tumor necrosis factor-alfa (TNF-α) of collagen-based biomaterials did not enhance the foreign-body reaction after implantation in mice (Khouw et al., 2001). CD44 (Bonnema et al., 2003), IFN-γ (Khouw et al., 1998, Khouw et al., 2000c), toll-like receptor 4 (TLR4) (Grandjean-Laquerriere et al., 2007), interleukin 4 (IL-4) (McNally and Anderson, 1995; DeFife et al., 1997), and nitric oxide (Hetrick et al., 2007) are some examples of host molecules involved in one way or another in the onset of the foreign-body reaction that can act as important targets for modulating the reaction to natural-based biomaterials. The encapsulation of nitric oxide and posterior release at the implantation site decreased the capsule formation around the implant and induced vascularization of the injured area (Hetrick et al., 2007). It was recently proved that TLR4 is involved in the release of TNF-α by natural-based biomaterials-activated macrophages (Grandjean-Laquerriere et al., 2007), although deeper knowledge regarding the involved pathway is still missing. It was also shown that the foreign-body reaction to collagenbased biomaterials can be delayed with a local injection of IFN-γ (Khouw et al., 1998), but treatment with the same cytokine in a systemic approach was shown to increase the cellular ingrowth and degradation of the biomaterial (Khouw et al., 2000c). Nonetheless, it was proven that IFN-γ inhibits the expression of the major histocompatibility complex (MHC) class II antigen by infiltrating cells into the biomaterial (Khouw et al., 1998; Khouw et al., 2000c). Khouw and co-workers showed that, either in mice or in rats, IFN-γ was not essential for the fusion of macrophages into foreign-body giant cells after the implantation of natural-based biomaterials (Khouw et al., 1998; Khouw et al., 2000b; Khouw et al., 2000c). However, it was demonstrated that, in humans, IL-4 is a potent macrophage fusion factor (McNally and Anderson, 1995; DeFife et al., 1997), contributing to the development of foreign-body reactions. CD44 was also found to be an important molecule in the modulation of the FBGCs formation (Bonnema et al., 2003). Usually, this type of modulation, involving the host cells and molecules, is carried out in particular cases where the developed biomaterials or TE devices already accomplish adequate physicochemical characteristics for the desired function. In addition, it is aimed to induce host interaction with the implanted biomaterial or device, and not to provoke major modifications in the immune system of the host, which may compromise the general physiological status and the performance of the biomaterial. Despite the work that has been carried out, further studies are needed to generate knowledge to modulate undesirable reactions to natural-based biomaterials.

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Final remarks

Great efforts have been done by research groups in biomaterials, biotechnology and biomedicine fields all over the world to understand the host tissue responses to the implantation of natural-origin biomaterials. Despite all this research, several mechanisms, interactions and factors involved in such responses remain to be identified and to be totally understood, which mainly arises from the enormous complexity of the tissues, the behaviour of different cells, and from the refined mammal immune system. Therefore, much basic investigation has still to be performed to solve or, at least, to try to fill the lacunae between the biomaterials and the immunology fields. Several approaches attempting to control the tissue response to the implantation of newly developed biomaterials have been tried, and despite the recent progress, ideal strategies, either concerning the material or the host, remain to be achieved. Furthermore, the defined approaches will always depend on the type of materials, on their physicochemical characteristics and specific function in the host, and on the physiological/pathological characteristics of the host itself.

26.7

Acknowledgements

This work was partially supported by the European Union funded STREP Project HIPPOCRATES (NMP3-CT-2003-505758) and the European NoE EXPERTISSUES (NMP3-CT-2004-500283).

26.8

References

Ali S A M, Doherty P J and Williams D F (1994), The mechanisms of oxidative degradation of biomedical polymers by free radicals, Journal of Applied Polymer Science, 51, 1389–98. Anderson J M (2000), Multinucleated giant cells, Curr Opin Hematol, 7, 40–7. Atkinson J P and Farries T (1987), Separation of self from non-self in the complement system Immunol Today, 8, 212–15 Bellingan G J, Caldwell H, Howie S E, Dransfield I and Haslett C (1996), In vivo fate of the inflammatory macrophage during the resolution of inflammation: inflammatory macrophages do not die locally, but emigrate to the draining lymph nodes, J Immunol, 157, 2577–85. Billingham M (1992), Normal Heart. In Sternberg S (Ed.) Histology for Pathologists, New York, Raven Press. Bonizzi G and Karin M (2004), The two NF-kappaB activation pathways and their role in innate and adaptive immunity, Trends Immunol, 25, 280–8. Bonnema H, Popa E R, Van Timmeren M M, Van Wachem P B, De Leij L F and Van Luyn M J (2003), Distribution patterns of the membrane glycoprotein CD44 during the foreign-body reaction to a degradable biomaterial in rats and mice, J Biomed Mater Res A, 64, 502–8.

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Brodbeck W G, Voskerician G, Ziats N P, Nakayama Y, Matsuda T and Anderson J M (2003), In vivo leukocyte cytokine mRNA responses to biomaterials are dependent on surface chemistry, J Biomed Mater Res A, 64, 320–9. Ciardelli G and Chiono V (2006), Materials for peripheral nerve regeneration, Macromol Biosci, 6, 13–26. Defife K M, Jenney C R, Mcnally A K, Colton E and Anderson J M (1997), Interleukin13 induces human monocyte/macrophage fusion and macrophage mannose receptor expression, J Immunol, 158, 3385–90. Fantone J and Ward P (1999), Inflammation. In Rubin E and Farber J (Eds.) Pathology, Third edition ed., Philadelphia P A, Lippincott-Raven. Goldsby R A, Kindt T J and Osborne B A (2000), Kuby Immunology, USA, W H Freeman and Company. Gorbet M B and Sefton M V (2004), Biomaterial-associated thrombosis: roles of coagulation factors, complement, platelets and leukocytes, Biomaterials, 25, 5681–703. Grandjean-Laquerriere A, Tabary O, Jacquot J, Richard D, Frayssinet P, Guenounou M, Laurent-Maquin D, Laquerriere P and Gangloff S (2007), Involvement of toll-like receptor 4 in the inflammatory reaction induced by hydroxyapatite particles, Biomaterials, 28, 400–4. Griffiths M M, Langone J J and Lightfoote M M (1996), Biomaterials and Granulomas, Methods, 9, 295–304. Harrist T, Schapiro B, Quinn T and Clark W (1999), The Skin. In Rubin E and Farber J (Eds.) Pathology, Third edition ed., Philadelphia P A, Lippincott-Raven. Hays A and Armbrustmacher V (1999), Skeletal Muscle. In Rubin E and Farber J (Eds.) Pathology, Third edition ed., Philadelphia P A, Lippincott-Raven. Heffner J R R (1992), Skeletal Muscle. In Sternberg S (Ed.) Histology for Pathologists, New York, Raven Press. Hetrick E M, Prichard H L, Klitzman B and Schoenfisch M H (2007), Reduced foreign body response at nitric oxide-releasing subcutaneous implants, Biomaterials, 28, 4571– 80. Ho T Y, Chen Y S and Hsiang C Y (2007), Noninvasive nuclear factor-kappaB bioluminescence imaging for the assessment of host-biomaterial interaction in transgenic mice, Biomaterials, 28, 4370–7. Hu W J, Eaton J W, Ugarova T P and Tang L (2001), Molecular basis of biomaterialmediated foreign body reactions, Blood, 98, 1231–8. Hunt J A (2001), Inflammation in Buschow K H J, Cahn R, Flemings M C and Ilscher B, Encyclopedia of Materials: Science and Technology, Oxford, Elsevier Science Ltd. Hunt J A and Williams D F (1995), Quantifying the soft tissue response to implanted materials, Biomaterials, 16, 167–70. Hunt J A, McLaughlin P J and Flanagan B F (1997), Techniques to investigate cellular and molecular interactions in the host response to implanted biomaterials, Biomaterials, 18, 1449–59. Johnson K, Chensue S and Ward P (1999), Immunopathology. In Rubin E and Farber J (Eds.) Pathology, Third edition ed., Philadelphia P A, Lippincott-Raven. Junqueira L C and Carneiro J (2005), Basic Histology: Text & Atlas, New York, McGrawHill. Kao W J and Lee D (2001), In vivo modulation of host response and macrophage behavior by polymer networks grafted with fibronectin-derived biomimetic oligopeptides: the role of RGD and PHSRN domains, Biomaterials, 22, 2901–9. Khouw I M, Van Wachem P B, De Leij L F and Van Luyn M J (1998), Inhibition of the

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tissue reaction to a biodegradable biomaterial by monoclonal antibodies to IFN-gamma, J Biomed Mater Res, 41, 202–10. Khouw I M, Van Wachem P B, Molema G, Plantinga J A, De Leij L F and Van Luyn M J (2000a), The foreign body reaction to a biodegradable biomaterial differs between rats and mice, J Biomed Mater Res, 52, 439–46. Khouw I M, Van Wachem P B, Plantinga J A, De Leij L F and Van Luyn M J (2001), Enzyme and cytokine effects on the impaired onset of the murine foreign-body reaction to dermal sheep collagen, J Biomed Mater Res, 54, 234–40. Khouw I M, Van Wachem P B, Plantinga J A, Haagmans B L, De Leij L F and Van Luyn M J (2000b), Foreign-body reaction to dermal sheep collagen in interferon-gammareceptor knock-out mice, J Biomed Mater Res, 50, 259–66. Khouw I M, Van Wachem P B, Van Der Worp R J, Van Den Berg T K De Leij L F and Van Luyn M J (2000c), Systemic anti-IFN-gamma treatment and role of macrophage subsets in the foreign body reaction to dermal sheep collagen in rats, J Biomed Mater Res, 49, 297–304. Liao H, Mutvei H, Sjostrom M, Hammarstrom L and LI J (2000), Tissue responses to natural aragonite (Margaritifera shell) implants in vivo, Biomaterials, 21, 457–68. Lickorish D, Chan J, Song J and Davies J E (2004), An in-vivo model to interrogate the transition from acute to chronic inflammation, Eur Cell Mater, 8, 12–9, discussion 20. Luttikhuizen D T, Harmsen M C and Van Luyn M J (2006), Cellular and molecular dynamics in the foreign body reaction, Tissue Eng, 12, 1955–70. Marques A P, Reis R L and Hunt J A (2005), An in vivo study of the host response to starch-based polymers and composites subcutaneously implanted in rats, Macromol Biosci, 5, 775–85. Mcnally A K and Anderson J M (1995), Interleukin-4 induces foreign body giant cells from human monocytes/macrophages. Differential lymphokine regulation of macrophage fusion leads to morphological variants of multinucleated giant cells, Am J Pathol, 147, 1487–99. Meinel L, Hofmann S, Karageorgiou V, Kirker-Head C, McCool J, Gronowicz G, Zichner L, Langer R, Vunjak-Novakovic G and Kaplan D L (2005), The inflammatory responses to silk films in vitro and in vivo, Biomaterials, 26, 147–55. Metz M and Maurer M (2007), Mast cells – key effector cells in immune responses, Trends Immunol, 28, 234–41. Mikos A G, McLntire L V, Anderson J M and Babensee J E (1998), Host response to tissue engineered devices, Adv Drug Deliv Rev, 33, 111–39. Mollnes T E (1997), Biocompatibility: complement as mediator of tissue damage and as indicator of incompatibility, Exp Clin Immunogenet, 14, 24–9. Mollnes T E (1998), Complement and biocompatibility, Vox Sang, 74 Suppl 2, 303–7. Nilsson B, Ekdahl K N, Mollnes T E and Lambris J D (2007), The role of complement in biomaterial-induced inflammation, Mol Immunol, 44, 82–94. Nimeri G, Ohman L, Elwing H, Wettero J and Bengtsson T (2002), The influence of plasma proteins and platelets on oxygen radical production and F-actin distribution in neutrophils adhering to polymer surfaces, Biomaterials, 23, 1785–95. Ratner B D (2007), The catastrophe revisited: blood compatibility in the 21st century, Biomaterials, doi:10.1016/j.biomaterials.2007.07.035. Ravin A G, Olbrich K C, Levin L S, Usala A L and Klitzman B (2001), Long- and shortterm effects of biological hydrogels on capsule microvascular density around implants in rats, J Biomed Mater Res, 58, 313–8. Robitaille R, Dusseault J, Henley N, Desbiens K, Labrecque N and Halle J P (2005),

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Inflammatory response to peritoneal implantation of alginate-poly-L-lysine microcapsules, Biomaterials, 26, 4119–27. Ruelicke T, Montagutelli X, Pintado B, Thon R and Hedrich H J (2007), Guidelines for the production and nomenclature of transgenic rodents – Report of the Felasa working group, Felasa. Spargo B J, Rudolph A S and Rollwagen F M (1994), Recruitment of tissue resident cells to hydrogel composites: in vivo response to implant materials, Biomaterials, 15, 853– 8. Stevens A, Lowe J S and Young B (2002), Wheater’s Basic Histopathology: A Colour Atlas and Text, Edinburgh, Churchill Livingstone. Szaba F M and Smiley S T (2002), Roles for thrombin and fibrin(ogen) in cytokine/ chemokine production and macrophage adhesion in vivo, Blood, 99, 1053–9. Tang L, Ugarova T P, Plow E F and Eaton J W (1996), Molecular determinants of acute inflammatory responses to biomaterials, J Clin Invest, 97, 1329–34. Urmacher C (1992), Normal Skin. In Sternberg S (Ed.) Histology for Pathologists, New York, Raven Press. Usami Y, Okamoto Y, Takayama T, Shigemasa Y and Minami S (1998), Chitin and chitosan stimulate canine polymorphonuclear cells to release leukotriene B4 and prostaglandin E2, J Biomed Mater Res, 42, 517–22. Van Luyn M J, Khouw I M, Van Wachem P B, Blaauw E H and Werkmeister J A (1998), Modulation of the tissue reaction to biomaterials, II. The function of T cells in the inflammatory reaction to crosslinked collagen implanted in T-cell-deficient rats, J Biomed Mater Res, 39, 398–406. Van Luyn M J, Plantinga J A, Brouwer L A, Khouw I M, De Leij L F and Van Wachem P B (2001), Repetitive subcutaneous implantation of different types of (biodegradable) biomaterials alters the foreign body reaction, Biomaterials, 22, 1385–91. Van Luyn M J, Van Wachem P B, Leta R, Blaauw E H and Nieuwenhuis P (1994), Modulation of the tissue reaction to biomaterials, I. Biocompatibility of cross-linked dermal sheep collagens after macrophage depletion, J Mater Sci: Mater Med, 5, 671– 678. Wheater P, Burkitt H and Daniels V (1979), Functional Histology: A Text and Colour Atlas, Edinburgh, Churchill Livingstone. Williams D F (2001), Biocompatibility principles. In Buschow K H J, Cahn R, Flemings M C and Ilsher B, Encyclopedia of Materials: Science and Technology, Oxford Elsevier Science Ltd.

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27 Immunological issues in tissue engineering N. R O T T E R, Ulm University, Germany

27.1

Introduction

Due to the shortage of donor organs and due to the increase in chronic diseases in combination with the increasing aging of the population the need for tissues and organs to repair damaged tissue function is continuously growing in all medical areas. In this context tissue engineering holds great promise for the regeneration of tissues and organs. Based on an interdisciplinary research effort it combines principles of engineering, biology and medical sciences. Currently different strategies for various applications are the subject of extensive investigation. A cell based strategy, sometimes termed in vitro tissue engineering is to isolate cells and culture them in vitro on threedimensional resorbable biomaterials with or without the addition of growth factors. Cells in this context might be chemically or genetically engineered in vitro in order to improve specific metabolical or mechanical functions. Another strategy is to use specific biomaterials in vivo in order to attract the host’s own cells to the damage site without prior in vitro cell culture. A cascade of immune reactions is initiated starting with the surgical procedure to implant the engineered tissue or biomaterial and continuing with biomaterial degradation and increasing tissue function. It significantly influences the implanted tissues structure and function. Until today many mechanisms and especially the molecular basis of these reactions remain incompletely elucidated.1 This chapter focuses on general issues of hosttransplant-reactions and gives examples and perspectives of strategies to prevent or benefit from these reactions in the context of tissue engineering.

27.2

Immune reactions to biomaterials

Biomaterials and in vitro engineered cells and tissues are introduced into the body by open surgery, endoscopically or by injection. These surgical measures primarily cause tissue trauma in various degrees. Additionally it is well known that a variety of adverse tissue reactions like inflammation, fibrosis, 699 © 2008, Woodhead Publishing Limited

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infection and thrombosis2 are triggered by biomaterials. Inflammation is dominated by activation of cells of the circulating blood, local inflammatory cells and endothelial cells of the microcirculation in tissue adjacent to the implant site.3 In association with the activation of blood plasma activator systems the reactions vary with regard to the implanted material properties like size, shape, surface characteristics, biodegradability and mechanical properties.4 Immediately after implantation biomaterials are covered by a protein layer.2 This protein layer is likely to be an integral part in controlling the host’s reactions to the biomaterial, as the host cells primarily interact with the adsorbed proteins and not with the material itself. It was demonstrated that albumin, fibrinogen and immunoglobulins are important parts of the protein layer5 and that conformational changes occur upon contact with the biomaterial surface which may lead to the exposure of hidden epitopes, which in turn might initiate the inflammatory response. The mechanisms of how this response takes place in detail are still unknown; however, it became clear that fibrinogen is not only essential in initiating the inflammatory response but also to the fibrous reaction.2 In the early phase after implantation the implant is surrounded by a fluid space containing cells and proteins.6 Polymorphonuclear leukocytes, monocytes and lymphocytes form the majority of cells in the first one to two weeks being an integral part of foreign body responses to biomaterials. Depending on the biomaterial’s surface properties macrophages and foreign body giant cells will persist at the site of the implant for a long time causing chronic inflammatory responses. At the same time regenerative responses are initiated with injured cells regenerating, angiogenesis taking place, matrix neosynthesis and remodelling occurring. Programmed cell death is limiting the regenerating cell populations.3 These reactions in turn lead to fibrous encapsulation of the implanted material.3 The fibrotic tissue is composed of fibroblasts and collagen.7 While the formation of thick fibrotic capsules contributes to the failure of joint implants8 and eye implants9 as well as encapsulated cells,10 poor integration of fibrotic tissue with artifical tissue might at the same time cause implant failure.11 In this complex stage of cellular reactions macrophages play an integral role in mediating the first adherence to the biomaterial. Their fusion into foreign body giant cells is the initial step of the chronic foreign body response.12 However the exact role of foreign body giant cells has still not been elucidated. The role of lymphocytes in these processes has been clarified further recently.12 There is evidence that lymphocytes are recruited and activated by cytokines which are produced following monocyte adhesion to the biomaterial. Most likely IL-2 and IL-6 are important molecules for the activation of lymphocytes,12,13 while the latter are thought to release IL-4 and IL-13, which are known to induce fusion of macrophages into foreign body giant cells.14,15 Dendritic cells are antigen-presenting cells connecting innate immunity including inflammation and adaptive immunity.16,17 This fact makes them

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important candidates in the immunity of tissue engineered constructs where the different immune responses are combined due to the cell-biomaterialconfiguration of the engineered tissues. Dendritic cells are activated via a family of receptors called Toll-like receptors (TLRs). Interestingly it has been shown that PLGA induces the maturation of dendritic cells18,19 and stimulates the secretion of proinflammatory cytokines like TNF-α and IL-6.19 Other materials like hyaluronic acid were demonstrated to modulate inflammatory chemokines and receptors as well as catabolic and inhibiting factors like MMPs/TIMP in human mesenchymal stem cells in vitro.20 However most in vivo studies were unable to demonstrate similar effects in vivo. This discrepancy is most likely due to the increased cell surface interactions taking place in the organism which involve the whole range of cells and mediators in a complex system. A deeper insight into the mechanisms of cell-biomaterial interactions concerning resident inflammatory cells as well as in vitro seeded cells is of utmost concern for tailoring biomaterials with respect to inflammatory and innate immune reactions.

27.3

Host reactions related to the implant site

In patients with osteoarthritis and rheumatoid arthritis high levels of fibronectin fragments are found in the synovial fluid.21,22 As these fibronectin fragments are potent inducers of proinflammatory cytokines and chemokines which can be produced by chondrocytes themselves it becomes clear that the success of tissue engineered cartilage in diseases in which high levels of fibronectin fragments are present will strongly depend on the suppression of the signaling pathways activated by fibronectin fragments. It has been demonstrated that fibronectin fragments21 as well as proinflammatory cytokines like IL-1β and TNF-α23,24 stimulate the expression of proinflammatory cytokines IL-6, IL8, GRO-α, GRO-β, and GRO-γ as well as the chemokine MCP-1 in chondrocytes. The use of tissue engineered cartilage in rheumatoid arthritis and osteoarthritis therefore can only be successful when the underlying pathological conditions which induce chondrocytic chondrolysis22 are blocked at the same time to a sufficient amount. Otherwise the surgically inserted engineered tissues will be exposed to the mentioned proinflammatory cytokines and will inevitably undergo the same fate of chondrolysis as the native tissue.

27.4

Immune reactions to different types of cells

27.4.1 Autologous cells and transplants In many tissue engineering applications autologous cells are used as a basis to engineer tissue for an individual patient. Generally autologous cells should

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not evoke an immune response at all. However isolation of cells from mature or immature tissue and their in vitro culture in monolayers and on biomaterials induces changes of cell surface characteristics. Especially chondrocytes naturally embedded in a dense matrix composed of collagen and glycosaminoglycans are not exposed to the immune system under physiological conditions. There is experimental evidence that monolayer culture leads to the exposure and change of cell surface antigens which then might lead to a significant immunological response.1,25 Cell isolation and amplification For in vitro tissue engineering procedures, small numbers of cells, in the case of cartilage tissue engineering, chondrocytes, are obtained by minimally invasive biopsies. These chondrocytes require significant amplification in monolayer culture in order to obtain sufficient cell numbers. During monolayer however the cells undergo significant changes in phenotype, surface characteristics and functionality. They not only switch from a round cell type to a more elongated fibroblast-like type, they furthermore stop the production of collagen type II and start to produce collagen type I instead.26 Also, the surface characteristics change with ongoing culture time, as was demonstrated by flow cytometry.27 The expression of ICAM-1 was demonstrated on human nasal chondrocytes following monolayer culture.25 Taking into account that chondrocytes are embedded within a dense cartilaginous matrix under physiological conditions and that significant changes occur during cell culture procedures, it becomes clear that even in an autologous setting immunological reactions can be directed against autologous cells, therefore possibly enhancing inflammatory reactions directed against biomaterials used as cell carriers. However until now it has not been elucidated which parts of the inflammatory reaction can be attributed to the cells and which parts are directed against the biomaterials. Further detailed investigations are required to distinguish these main factors and to target them therapeutically.

27.4.2 Allogeneic cells and transplants As autologous cells might not be available in certain patients, for example, due to extensive trauma and pre-existing surgical defects in the case of cartilage tissue engineering or due to the patient’s age, in endothelialization of artifical prothesis in cardiovascular bypass surgery28 the use of allogeneic cells could be a valuable option in these specific cases. However allogeneic cells carry the risk of immunological rejection due to mismatches in the Major Histocompatibility Complex (MHC) system. The immunoresponse to alloantigens can be cell-mediated or humoral. Cell mediated reactions seem to play a more pronounced role, but antibodies

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might contribute to the response. The antigens exposed in the graft stimulate responses of the host’s immune system with different types of lymphocytes responding. Antigen presenting cells recognize the transplanted allogeneic cells, take up and process antigens of these cells and trigger the rejection, which is composed of cytotoxic T cells, allospecific antibodies and complement.29,30 The extent and the time of such a rejection depend on different factors, like whether the transplant is primarily or secondarily vascularized, whether it is connected to the lymph vessels and whether it was physically or chemically pre-treated. Furthermore it depends on the immunsuppression of the host.31 Genetic modification of allogeneic cells could be a promising strategy to reduce cellular and humoral allorejection mechanisms. It was demonstrated by Doebis et al.28 that down-regulated MHC class I expression in endothelial cells of the rat using an intracellularly expressed antibody directed against MHC class I molecules reduces allorejection mechanisms. Furthermore some authors were even able to demonstrate that the transplantation of allogeneic tissue-engineered endothelial cell constructs may provide long-term control of vascular repair after injury.32 There is also evidence that endothelial cells embedded in three-dimensional matrix constructs consisting of gelatine express reduced levels of MHC class II molecules as compared to endothelial cell suspensions or endothelial cells grown on tissue culture plastic.33 These experiments provide first evidence that allogeneic cells might be used without additional modifications of their immunogeneity. Mesenchymal stem cells (MSCs) might also be derived from allogeneic donors. They can be derived from various tissue sources34 and are a potential alternative source to chondrocytes and other differentiated cells in tissue engineering applications due to their expansion and differentiation potential.35 As surface characteristics include the absence of immunologically important markers like HLA-DR (MHC-II), CD 40, CD-40-ligand, CD80 and CD86 and immunomodulatory effects have been demonstrated in vitro37 it has been proposed that they might be a valuable source to be used without additional immunosuppressive agents. It has been demonstrated that undifferentiated mesenchymal stem cells do not evoke a stimulation of allogeneic lymphocytes in vitro, they rather seem to suppress stimulated allogeneic lymphocytes.38 Also osteogenic differentiation in vitro, for example, does not seem to inhibit these effects.39 However in vivo, and especially following induction of differentiation with concurrent changes in cell function and additional antigenicity, for example, gene-modifying strategies will have to be applied to ensure transplant survival.40 Alternative stem cell sources have been explored recently with regard to immunomodulatory properties in vitro. Wollbank et al. demonstrated that mesenchymal and epithelial amniotic stem cell populations, and adipose tissue-derived stem cells, exhibit time, contact- and dose-dependent immunosuppressive characteristics in vitro and might therefore

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be another alternative for allogeneic applications in tissue engineering.41 Also the use of MSCs as immunosuppressant agents in autoimmune diseases has been proposed and successfully tested in animal models. Augello et al. demonstrated that one injection of MSCs prevented the occurrence of experimental collagen induced arthritis.42 However there are contradictory studies43 which did not show any benefit of the injection of stem cells in the same experimental model. In summary it currently remains unclear whether the application of allogeneic stem cells might be a valuable alternative to autologous cells. Further and more detailed analyses in vitro and in vivo are required to finally evaluate the conditions of the applicability of these cells in an allogeneic setting. Of course there would be enormous benefits by this ‘off the shelf’ availability of engineered cells and tissues. They not only would reduce therapy time but also medical concerns when seeking for alternative treatment options.

27.4.3 Xenogeneic cells and transplants The transplantation of xenogenic tissues and cells offers the unexcelled advantage of unlimited availability on the one hand but raises ethical, immunological and infectious concerns on the other hand. Although there have been major advances in the field of xenotransplantation like the production of GalT-KO pigs showing that the problem of hyperacute rejection can be solved,44 many other issues like acute vascular rejection and cellular xenograft rejection have not yet been completely elucidated. Preclinical studies, such as Hering et al.’s who transplanted porcine pancreatic islets into immunosuppressed diabetic monkeys, demonstrated that specific immunosuppressive regimens can lead to normoglycemia for up to 100 days.45 However, the problems of xenozoonoses, like PERV (porcine endogenous retrovirus) infections, remains unsolved.46 In the context of tissue engineering when autologous or allogeneic cells are available the transplantation of xenogeneic cells therefore currently seems an unfavorable option. However, the use of decellularized xenogeneic tissues have potential, e.g. for heart valve replacement or vocal fold repair. Although detailed analyses of decellularized xenogeneic tissues are still missing, the general principles mentioned in the materials part apply.

27.5

Immune reactions to in vitro engineered tissues

An in vitro engineered tissue is composed of a cellular and a material component usually associated with matrix proteins which were produced during in vitro cell culture. The transplantation of cells and matrix proteins leads to an antigen-specific immune response47 while the biomaterial component induces

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a non-specific inflammatory and regenerative response. Protein adsorption, complement activation, coagulation and the adhesion and activation of macrophages play important roles in this part of the reaction, while there is evidence that polymeric carriers without associated antigens do not evoke an antigen-specific immune reaction.48 Antigens of the cellular or protein component in engineered tissues are processed and presented jointly with the host MHC class II molecules by antigen presenting cells like macrophages or dendritic cells. Some studies have been carried out to determine the relationship and the interdependence between the immune responses to the cellular component and the inflammatory responses to the biomaterials component. It has been demonstrated that the biomaterial component which induces an inflammatory response with the recruitment of various cell types amplifies the humoral response to the associated cellular and protein antigens47 by using the model antigen ovalbumin. In this context the structure of the biomaterial has a crucial influence on the intensity of the reaction as microparticles only temporarily and moderately induced the humoral immune response while scaffolds induced a stronger and permanent response, although the amount of polymer and antigen delivered was identical. It was hypothesized that the way of surgical insertion (injection versus open surgery) was causing the differences in immune response, because open surgery causes significantly enhanced tissue damage with necrosis and inflammation while injection is minimally invasive with regard to local tissue damage. Necrosis and inflammation can be considered ‘danger signals’ which are indispensable prerequisites for an immune response according to the ‘danger hypothesis’. The total surface area as a crucial factor in the difference of the humoral immune response seems unlikely as it was calculated to be almost similar in the cited study.

27.6

Immune protection of engineered constructs

Considering the different factors involved in the implantation of tissue engineered constructs there is a large variety of different possible strategies to reduce inflammation, resorption and rejection of these tissues. As discussed before, the mode of applying the engineered tissue – minimally invasive versus open surgery – is of utmost concern as strong inflammatory reactions can be induced by open surgery alone.47 This leads to the conclusion that minimizing the surgical trauma is the first and basic step in immunoprotection of an engineered tissue. On the other hand some procedures like the implantation of large volume engineered tissues will always require open surgical approaches. In these cases the local control of the surgically induced inflammatory reaction will be essential. Other possible strategies involve modifications of the biomaterial or the in vitro engineered cells, which apply to the engineered tissues themselves

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while other strategies target components of the innate and adaptive immune system of the host by applying therapeutic agents to the host organism directly.

27.7

Strategies directed towards reactions to biomaterials

Modifications of the biomaterials concerning composition, structure, surface area and degradation time play a key role in tailoring the host’s response to the implanted tissue. However the limited knowledge on biomaterial–cell interactions2 in vivo currently disables an ‘off-the-shelf’ design of the optimal biomaterial for a specific application and still requires a whole range of evaluation methods including in vitro and in vivo strategies. Strategies concerning material production and modification are discussed in the other chapters. The local application of anti-inflammatory agents could of course reduce the acute inflammatory reaction related to the implantation of the biomaterial. The immobilization of urokinase on polyurethane tubes, for example, significantly reduced the acute inflammatory reaction.49 However the fibrotic reaction remained unchanged by this treatment. The dispersion of non-steroidal antiinflammatory drugs (NSAIDs) into biodegradable polymeric matrices has been proposed as a way to obtain good therapeutic effects in joints while minimizing the side effects of NSAIDs.50 This approach along with the use of glucocorticoids51 might be useful in tissue engineering applications as well. Steroids have among others a strong anti-inflammatory effect. This is due to reduction of activation and chemotaxis of inflammatory cells and due to modifications in the arachidonic acid metabolism. Furthermore they down regulate capillary dilatation and suppress the expression of adhesion molecules and cytokines. They also suppress the influx of macrophages and neutrophils which are known to play an essential role in the initial inflammatory response to the implantation of the biomaterial. Depending on the application, the use of such agents could however interfere with the development of the engineered tissue. Haisch et al. demonstrated the growth of bone instead of cartilage by methylprednisolonestimulated chondrocytes51 in a rabbit model, although the local inflammatory reaction was significantly reduced by the addition of methylprednisolone. Furthermore the systemic application of such potent drugs will dosedependently lead to side effects related to the multitude of effects. On the other hand a significant inflammation might be beneficial in implantation sites with reduced vascularization, e.g. following radiation therapy. The degree and length of inflammation therefore should depend on the application and implantation site and should ideally be tailored specifically for a distinct implant and implantation site.

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Strategies directed towards reactions to implanted cells

So far it is not known in detail to what extent cell isolation and amplification in vitro modify antigenic properties of autologous cells.1 However it has been demonstrated that autologous chondrocytes isolated from the cartilaginous matrix express ICAM-1. This enables them to elude an immune response, although being autologous cells.25 On the other hand differentiated cells like chondrocytes have been used successfully in clinical applications,52,53 making relevant changes in their antigen structure unlikely. However in the head and neck, engineered tissues have only been used occasionally54–57 and are not routinely used. These differences might be due to variations in site specific immunological reactions, which make the subcutaneous localization of the head and neck more prone to immune reactions as compared with the joint. Strategies which could be followed in the use of autologous cells when necessary as well as in non-autologous cells are immunoisolation, local immunosuppression and genetic modification to reduce the adaptive immune response.

27.8.1 Immunoisolation Immunoisolation can be achieved by the encapsulation of cells with biomaterials that prevent the cells from being exposed to the immune system on the one hand but allow diffusion of nutrients and metabolic factors on the other hand. Encapsulation has been applied in the transplantation of central nervous system cells58 and β-cells for the treatment of insulin-dependent diabetes mellitus59 but it has also been used in cartilage tissue engineering.60–62 However it has to be ensured that the material used for encapsulation, e.g. highlypurified alginate, is stable over a long time period63,64 and that it does not evoke a strong inflammatory response itself, as this might lead to the degradation of the encapsulation with concurrent destruction of the transplant.

27.8.2 Immunomodulation Genetic engineering of implanted cells Immunomodulation which means in most applications immunosuppression might be induced by genetical engineering of the donor cells in a way that MHC and other molecules are repressed leading to the suppression or absence of an antigen-specific immune response. Depletion of MHC class I expression on the surface of endothelial cells in a rat model was achieved by the application of an intracellularly expressed antibody directed against MHC class I molecules.28 This resulted in a protection of these cells against killing by allo-specific cytotoxic T cells and antibody-complement mediated lysis. Other

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authors modified immortalized rat hepatocytes by transduction with products of early region 3 (AdE3) of the adenoviral genome, as they are known to protect infected cells from immune recognition and destruction. They demonstrated that this measure can protect these cells from allorejection.65 These results demonstrate that genetic engineering of cells in vitro is a promising pathway for future applications in tissue engineering. Local and systemic pharmacological immunomodulation Blocking T- and B-cell response could be suitable in certain applications to suppress the immune response to the tissue engineered construct. Immunosuppressive agents are widely used in allotransplantation. However their systemic application is associated with a wide spectrum of severe side effects such as an increased rate of severe bacterial, viral and fungal infections,66 and increased formation of malignancies.67 In the context of tissue engineering a systemic application is of course an option in case of life-saving cell therapy; however, more frequently tissue engineers are aiming at the construction of tissues which make this type of treatment unnecessary. To reduce side effects, immunosuppressive agents can also be administered locally, as was demonstrated in a prolonged survival of skin allografts in mice68 by the administration of salen manganese complexes which are scavengers of reactive oxygen species and catalase activities on rejection and alloresponse. Tacrolimus and rapamycin, relatively new immunosuppressive agents have been cotransplanted with mesencephalic cells from fetal mice in the brain of rats, demonstrating better graft survival.69 Immunomodulation by monoclonal antibodies Modulating the immune response can also be performed more specifically than by general blocking of the B- and T-cell response, e.g. by the application of specific blocking antibodies. As macrophages play an integral role in the inflammatory and regenerative response to biomaterial implantation, modulating their function is one possibility to tailor tissue reactions to implants. Inhibition of macrophage recruitment can be induced by anti-inflammatory drugs applied locally like dexamethason (see Section 27.7). This can have certain disadvantages as was demonstrated in a rabbit model. Although the application led to a significantly reduced inflammatory reaction, it acted as an inducer of bone formation.51 Macrophage migration inhibitory factor (MIF) plays a key role in inflammation and immune responses. MIF is secreted by a variety of cells including macrophages70 and is upregulated by proinflammatory stimuli. MIF in turn is responsible for the production of proinflammatory cytokines like TNF-α, IL-1β and IL-671 while it activates T-cells and macrophages.

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MIF blocking antibodys have been, among others used successfully in animal models of collagen-induced arthritis72 or allergic airway inflammation.73 They, along with antibodies directed against other macrophage attracting chemokines might be used in tissue engineering applications to reduce the macrophage response. Dendritic cells could also be targeted to reduce immune responses to tissue engineered constructs as they have a key role in bridging innate and adaptive immunity.16,17 Blocking NF-κB with the inhibitor BAY 11-7082 specifically in dendritic cells was shown to induce antigen-specific immune suppression in experimental inflammatory arthritis.74 An unspecific inhibition of NF-κB could however result in adverse effects as the receptor is involved in the differentiation, activation, and survival of a multitude of cells. Tailored modulation of dendritic cells could however be an interesting option in tissue engineering. Depending on the type of cell implanted, other receptors might be specifically targeted in order to modify inflammatory as well as functional responses. One example is the α5β1 fibronectin receptor. Important signaling pathways, like the MAPK pathway in chondrocytes are activated through the α5β1 fibronectin receptor.75 This pathway in turn activates nuclear transcription factors regulating the gene expression of proinflammatory cytokines and chemokines. Direct inhibition of the α5β1 fibronectin receptor with blocking antibodies could be an option in modifying the immune response. This direct inhibition however was demonstrated to lead to a significant increase in cell death,76 making it not applicable in tissue engineering nor other therapeutic applications. Also it has to be taken into account that NF-κB plays a distinct role in normal immune processes, such as prevention of apoptosis in certain tissues. Therefore the inhibition of NF-κB may lead to harmful effects on the chondrocytes as well. Other options include the targeting of upstream molecules as IKK-2 or the inhibition of p38 or JNK MAPKs.21 This has been demonstrated to reduce matrix destruction in models of arthritis. The use of monoclonal antibodies directed against specific cellular components of immune reactions to the engineered tissues is most likely a promising option in tissue engineering applications because they do not lead to severe side effects nor do they interfere with the function of the implanted tissue.

27.9

Future trends

Tissue engineering and regenerative medicine hold great promise in many medical specialties. The therapy of chronic kidney, liver or heart failure, the treatment of diabetes and degenerative neurological diseases as well as the treatment of bone and joint diseases or complex tissue defects in the head and neck region might undergo revolutionary changes with tissue engineering entering the clinical realm.

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Currently allogeneic organ transplantation is the first treatment option for chronic organ failure. As these procedures evoke rejection responses of the host organism life long immunosuppression is inevitable. This in turn results in severe side effects and sometimes even then chronic rejection with dysfunction of the transplanted organ occurs. Furthermore there is a shortage of donor organs which makes thousands of transplant candidates wait for a long time and sometimes die during the waiting time for a donor organ. The in vitro growth of autologous or allogeneic cells in conjunction with resorbable biomaterials therefore seems to become a promising treatment option. With the increasing knowledge in genetic engineering and stem cell biology and with sophisticated biomaterial production technologies, the range of strategies to ensure survival of tissue engineered constructs is continuously growing. As mentioned in this text the variety of methods to ensure survival of engineered tissues has grown enormously. From the clinical point of view it is of utmost concern to develop strategies which are applicable in the clinic. This means that ideally a transplant should be available ‘off the shelf’ without a distinct waiting time which would however require the use of allogeneic or even xenogeneic cells. On the other hand the use of autologous cells is a great advantage as immunosuppression can be avoided. Although immunomodulatory strategies might be applied locally or by genetic modification of non-autologous cells their life long necessity leads to a more complex therapy with distinct side effects and makes the therapy significantly more expensive. Therefore autologous cells most likely will remain the cell source of choice for the majority of cases at least until tailoring and modification of allogeneic cells becomes a standard procedure and proves to be safe with regard to rejection and transmission of infectious diseases. Strategies to tailor the inflammatory and fibrous reaction to biomaterials with respect to a specific application have great potential. However the majority of factors involved and the way of interaction of these factors are currently largely unknown and therefore a distinct amount of basic research work will be necessary before these strategies can be applied in the clinic.

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54 Warnke P H, Springer I N, Wiltfang J, Acil Y, Eufinger H, Wehmoller M, Russo P A, Bolte H, Sherry E, Behrens E and Terheyden H (2004), Growth and transplantation of a custom vascularised bone graft in a man, Lancet, 364, 766–770. 55 Warnke P H, Wiltfang J, Springer I, Acil Y, Bolte H, Kosmahl M, Russo P A, Sherry E, Lutzen U, Wolfart S and Terheyden H (2006), Man as living bioreactor: fate of an exogenously prepared customized tissue-engineered mandible, Biomaterials, 27, 3163–3167. 56 Yanaga H, Yanaga K, Imai K, Koga M, Soejima C and Ohmori K (2006), Clinical application of cultured autologous human auricular chondrocytes with autologous serum for craniofacial or nasal augmentation and repair, Plast Reconstr Surg, 117, 2019–2030; discussion 2031–2032. 57 Yanaga H, Koga M, Imai K and Yanaga K (2004), Clinical application of biotechnically cultured autologous chondrocytes as novel graft material for nasal augmentation, Aesthetic Plast Surg, 28, 212–221. 58 Emerich D F, Winn S R, Christenson L, Palmatier M A, Gentile F T and Sanberg P R (1992), A novel approach to neural transplantation in Parkinson’s disease: use of polymer-encapsulated cell therapy, Neurosci Biobehav Rev, 16, 437–447. 59 Bloch K and Vardi P (2005), Toxin-based selection of insulin-producing cells with improved defense properties for islet cell transplantation, Diabetes Metab Res Rev, 21, 253–261. 60 Haisch A, Groger A, Gebert C, Leder K, Ebmeyer J, Sudhoff H, Jovanovic S, Sedlmaier B and Sittinger M (2005), Creating artificial perichondrium by polymer complex membrane macroencapsulation: immune protection and stabilization of subcutaneously transplanted tissue-engineered cartilage, Eur Arch Otorhinolaryngol, 262, 338–344. 61 Haisch A, Groger A, Radke C, Ebmeyer J, Sudhoff H, Grasnick G, Jahnke V, Burmester G R and Sittinger M (2000), Macroencapsulation of human cartilage implants: pilot study with polyelectrolyte complex membrane encapsulation, Biomaterials, 21, 1561– 1566. 62 Haisch A, Groger A, Radke C, Ebmeyer J, Sudhoff H, Grasnick G, Jahnke V, Burmester G R and Sittinger M (2000), [Protection of autogenous cartilage transplants from resorption using membrane encapsulation], HNO, 48, 119–124. 63 Thanos C G, Calafiore R, Basta G, Bintz B E, Bell W J, Hudak J, Vasconcellos A, Schneider P, Skinner S J, Geaney M, Tan P, Elliot R B, Tatnell M, Escobar L, Qian H, Mathiowitz E and Emerich D F (2007), Formulating the alginate-polyornithine biocapsule for prolonged stability: Evaluation of composition and manufacturing technique, J Biomed Mater Res A, 83A(1), 216–224. 64 Thanos C G, Bintz B E and Emerich D F (2007), Stability of alginate-polyornithine microcapsules is profoundly dependent on the site of transplantation, J Biomed Mater Res A, 81, 1–11. 65 Mashalova E V, Guha C, Roy-Chowdhury N, Liu L, Fox I J, Roy-Chowdhury J and Horwitz M S (2007), Prevention of hepatocyte allograft rejection in rats by transferring adenoviral early region 3 genes into donor cells, Hepatology, 45, 755–766. 66 Cronin D C 2nd, Faust T W, Brady L, Conjeevaram H, Jain S, Gupta P and Millis J M (2000), Modern immunosuppression, Clin Liver Dis, 4, 619–655, ix. 67 Jonas S, Rayes N, Neumann U, Neuhaus R, Bechstein W O, Guckelberger O, Tullius S G, Serke S and Neuhaus P (1997), De novo malignancies after liver transplantation using tacrolimus-based protocols or cyclosporine-based quadruple immunosuppression with an interleukin-2 receptor antibody or antithymocyte globulin, Cancer, 80, 1141– 1150. © 2008, Woodhead Publishing Limited

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68 Tocco G, Illigens B M, Malfroy B and Benichou G (2006), Prolongation of alloskin graft survival by catalytic scavengers of reactive oxygen species, Cell Immunol, 241, 59–65. 69 Alemdar A Y, Sadi D, McAlister V and Mendez I (2007), Intracerebral co– transplantation of liposomal tacrolimus improves xenograft survival and reduces graft rejection in the hemiparkinsonian rat, Neuroscience, 146, 213–224. 70 Calandra T, Bernhagen J, Mitchell R A and Bucala R (1994), The macrophage is an important and previously unrecognized source of macrophage migration inhibitory factor, J Exp Med, 179, 1895–1902. 71 Bernhagen J, Mitchell R A, Calandra T, Voelter W, Cerami A and Bucala R (1994), Purification, bioactivity, and secondary structure analysis of mouse and human macrophage migration inhibitory factor (MIF), Biochemistry, 33, 14144–14155. 72 Mikulowska A, Metz C N, Bucala R and Holmdahl R (1997), Macrophage migration inhibitory factor is involved in the pathogenesis of collagen type II-induced arthritis in mice, J Immunol, 158, 5514–5517. 73 Amano T, Nishihira J and Miki I (2007), Blockade of macrophage migration inhibitory factor (MIF) prevents the antigen-induced response in a murine model of allergic airway inflammation, Inflamm Res, 56, 24–31. 74 Martin E, Capini C, Duggan E, Lutzky V P, Stumbles P, Pettit A R, O’Sullivan B and Thomas R (2007), Antigen-specific suppression of established arthritis in mice by dendritic cells deficient in NF-kappaB, Arthritis Rheum, 56, 2255–2266. 75 Forsyth C B, Pulai J and Loeser R F (2002), Fibronectin fragments and blocking antibodies to alpha2beta1 and alpha5beta1 integrins stimulate mitogen-activated protein kinase signaling and increase collagenase 3 (matrix metalloproteinase 13) production by human articular chondrocytes, Arthritis Rheum, 46, 2368–2376. 76 Pulai J I, Del Carlo M Jr and Loeser R F (2002), The alpha5beta1 integrin provides matrix survival signals for normal and osteoarthritic human articular chondrocytes in vitro, Arthritis Rheum, 46, 1528–1535.

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28 Biocompatibility of hyaluronic acid: From cell recognition to therapeutic applications K. G H O S H, Children’s Hospital and Harvard Medical School, USA

28.1

Introduction

Hyaluronic acid (HA), also known as hyaluronan, is a ubiquitous, naturallyoccurring, polyanionic, glycosaminoglycan that consists of repeating nonsulfated disaccharide units (α-1,4-D-glucuronic acid and β-1,3-N-acetyl-Dglucosamine) of variable sizes, appearing in molecular weights ranging from 0.1 to 10 million Daltons (Fraser et al., 1997). HA was initially known to exhibit only unique physicochemical properties that help to maintain tissue viscoelasticity. However, subsequent studies revealed that HA also exerts important biological effects by binding to specific cell surface receptors and other extracellular matrix (ECM) molecules, which initiates intracellular signaling cascades that modulate key functions such as adhesion, migration and proliferation (Aruffo et al., 1990; Entwistle et al., 1996; Toole, 2004). Together, the complex biological and physicochemical properties of HA influence key developmental processes such as embryogenesis, morphogenesis and wound repair (Chen and Abatangelo, 1999; Toole, 2001). As a result, HA has attracted huge interest for use in various therapeutic applications (Vercruysse and Prestwich, 1998; Allison and Grande-Allen, 2006; Balazs and Denlinger, 1989). The purification of the non-inflammatory fraction of HA over three decades ago initiated a host of therapeutic trials that involved supplementation of unmodified HA into the site of defect (Balazs and Gibbs, 1970). Early results showed that HA was effective in protecting retinal damage during ophthalmic surgery, reducing wound scarring, preventing post-operative adhesions, and reducing pain while increasing mobility in arthritic joints (Denlinger and Balazs, 1980; Denlinger et al., 1980; Balazs and Denlinger, 1989); however, these effects were short-lived due to the rapid degradation of native HA by the HA-specific enzyme, hyaluronidase. To increase its residence time in vivo, HA has since been chemically modified and subsequently crosslinked using myriad approaches (Campoccia et al., 1998; Prestwich et al., 1998; Park et al., 2003). To further enhance their biological activity or produce 716 © 2008, Woodhead Publishing Limited

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tailor-made tissues, chemically-modified HAs have been derivatized with various ECM-derived peptides or protein fragments (Ghosh et al., 2006; Shu et al., 2004). Such HA derivatives have found great use as biomaterials in several medical applications such as drug delivery, wound repair and tissue engineering (Allison and Grande-Allen, 2006; Ghosh et al., 2006; Horn et al., 2007; Kim et al., 2007; Luo et al., 2000; Nettles et al., 2004; Shu et al., 2004). This chapter will discuss how the vast knowledge about the key role of HA in tissue development, homeostasis and repair has been leveraged to develop novel and potent therapeutic applications and highlight recent studies that implicate the use of HA in regenerative medicine.

28.2

Native hyaluronan

28.2.1 Occurrence Hyaluronan is found in various mammalian tissues across all vertebrates. In terms of net amount, almost half of the total HA per organism can be found in skin, with the musculo-skeletal system accounting for another quarter fraction of the total quantity. In terms of concentration per tissue, it is the highest in typical connective tissues such as synovial fluid and umbilical cord (~3 mg/ml), while the skin and vitreous humor (eye) also containing moderate concentrations of HA (~0.2 – 0.5 mg/ml). Detailed analyses of HA distribution in mammalian tissues has been published elsewhere (Reed et al., 1988; Laurent, 1981; Engstrom-Laurent et al., 1985; Tengblad et al., 1986; Laurent et al., 1996; Fraser et al., 1993; Laurent et al., 1995) and summarized by Fraser et al. (1997). HA can also be found in lung and kidney, while the lowest concentration has been reported in plasma. Interestingly, the highest concentration of HA in a mammalian tissue is found in rooster comb (7.5 mg/ml), which has long served as an important source for HA isolation (Manna et al., 1999; Swann, 1968; Swann and Caulfield, 1975). Importantly, the source of HA isolation should be carefully determined since HA obtained from different tissues and species contains varying amounts and types of contaminants that may alter its function both in vitro and in vivo (Shiedlin et al., 2004).

28.2.2 Biosynthesis HA is synthesized in the plasma membrane by a specialized enzyme called hyaluronan synthase, with the nascent chains being directly secreted into the extracellular space (Fraser et al., 1997; Prehm, 1983a; Watanabe and Yamaguchi, 1996). The enzyme alternately adds the sugar units from the activated nucleotide precursors (UDP – glucuronic acid and UDP-Nacetlyglucosamine) to the reducing end of the growing chain (Mian, 1986;

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Prehm, 1983b), which is in marked contrast with other glycosaminoglycans that grow at the non-reducing ends. These newly-synthesized HA chains can contain up to 10 000 repeat disaccharide units or more, with the molecular weights reaching up to and beyond four million Daltons. Previous studies have suggested that the final chain size is regulated by thermodynamic parameters, where the decrease in entropy during macromolecular synthesis is balanced by release of free energy during cleavage of repeat disaccharide units and subsequent chain organization (Nickel et al., 1998; Philipson et al., 1985). The secreted HA chains are very flexible and are usually found in the ECM in a randomly-coiled configuration; when stretched from end to end, these chains can extend up to ~10 µm. HA is synthesized by most tissue cells during their original life-span, although mesenchymal cells exhibit the strongest expression (Lee et al., 1993; Nishida et al., 1999; Evanko et al., 1999; Asplund et al., 1993; Heldin and Pertoft, 1993).

28.2.3 Physicochemical and structural properties HA is polyanionic at extracellular pH, which results from oxidation of the carboxylic group on HA. This property allows it to bind cations (e.g. Na+, Ca2+), leading to an increase in osmotic gradient that, in turn, attracts and binds water within the HA polymeric network (Laurent and Fraser, 1992; Comper and Laurent, 1978; Gribbon et al., 2000). As a result, the long HA chains swell and occupy enormous extracellular space. The bound water is largely immobilized, which causes steric exclusion by restricting free diffusion of fluids and other ECM molecules (Ogston and Sherman, 1961; Ogston and Phelps, 1961). In addition, despite being unipolar, HA chains interact with themselves through the creation of distinct hydrophobic patches along their backbones (Scott and Heatley, 1999; Mikelsaar and Scott, 1994). At higher concentrations, such chain-chain interactions form an entangled network, which confers to HA its unique viscoelastic property by its ability to resist (elastically) rapid, short duration fluid flow while undergoing partial realignment and viscous movement in response to show longer duration fluid flow (Furlan et al., 2005; Falcone et al., 2006). Such hygroscopic nature and unique biomechanical function of HA makes it an indispensable component of the vitreous humor and the ECM of cartilage and other skeletal joints (Weiss, 2000). In addition to its unique physicochemical functions, HA also provides important structural support to the ECM. Hyaluronan-binding proteins, called hyaladherins, mediate its interaction with various extracellular components, including proteoglycans, collagen and fibrin, that stabilizes both HA and the ECM (Toole, 2001; Chen et al., 1994). In cartilage, for example, the link protein promotes HA-aggrecan association that is crucial for HA stabilization and, together with its association with collagen fibrils, the resulting aggrecan-

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HA complex provides structural stability to the entire connective tissue (Nishida et al., 1999; Hardingham, 1981). Versican, an aggregating proteoglycan, also associates with HA and retains it in tissues through complex interactions involving fibronectin and collagen (Sorrell et al., 1999; Evanko et al., 1999). Another interesting manifestation of HA’s structural role is the formation of a pericellular coating seen (indirectly) around most cells of mesodermal origin through exclusion of cells and other particles (Lee et al., 1993; Knudson and Knudson, 1993; Knudson et al., 1993). Besides creating a local cellular microenvironment, the HA coating helps in fending off attacks by immune cells and viruses.

28.2.4 Biological function Role in inflammation In addition to exhibiting distinct physicochemical properties, HA also plays an important biological role through its ability to modulate inflammation following tissue injury, which imparts upon HA its superior biocompatibility. Tissue damage causes HA degradation into ‘active’ lower molecular weight (LMW) fragments, which can occur either through the action of hyaluronidase or due to non-enzymatic activities such as mechanical impact or free radical activity (Laurent and Fraser, 1992; Noble, 2002). The LMW HA activates pro-inflammatory cytokines and stimulates tissue cell proliferation, migration and angiogenesis that, collectively, promote tissue repair. Specifically, LMW HA causes toll-like receptor 4 (TLR 4)-mediated activation of dendritic cells and capillary endothelial cells, which secrete inflammatory cytokines such as tumor necrosis factor-α (TNF-α) and interleukin (IL)-1β, and IL-8 (Termeer et al., 2000; Termeer et al., 2002; Taylor et al., 2004). Importantly, this proinflammatory activity of HA is observed exclusively at LMW as improper degradation of HA leads to incomplete tissue repair (Termeer et al., 2000; Termeer et al., 2002; Noble, 2002). The inflammatory cytokines cause capillary endothelial cells to increase expression of HA, which interacts with the HA-specific CD44 receptors on lymphocytes to promote their recruitment to the site of inflammation (Siegelman et al., 1999; Mohamadzadeh et al., 1998). The CD44/HA interaction is a critical determinant of successful tissue repair; CD44-knockout (CD44-KO) mice are unable to repair bleomycin-induced lung damage and eventually die within two weeks (Teder et al., 2002). Detailed analyses showed that unlike wild-type mice, the CD44-KO mice had persistently high levels of inflammation and HA oligosaccharides, further supporting the role of LMW HA fragments in tissue inflammation. LMW HA fragments are also abundant in other pathologic conditions marked by chronic inflammation such as rheumatoid arthritis and chronic colon inflammation, among others (de la Motte et al., 2003; Laurent and Fraser, 1992; Laurent et al., 1995). © 2008, Woodhead Publishing Limited

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The higher MW (HMW) HA, on the other hand, exerts a contrasting effect on the reparative process by inhibiting both inflammation and angiogenesis (Day and de la Motte, 2005; Chen and Abatangelo, 1999). Recent data suggests that HMW HA chains form intermolecular crosslinks to create robust fibrils that organize into complex molecular scaffolds, which strongly bind leukocytes and sequester them from the underlying proinflammatory tissue cells, thereby limiting tissue inflammation (Day and de la Motte, 2005). This spontaneous crosslinking of HMW HA chains is mediated by four proteins, viz. inter-α-inhibitor (IαI), pre-α-inhibitor (PαI), pentraxin 3 (PTX3) and TSG-6 (a 35 kDa-secreted product of the tumor necrosis factor-stimulated gene-6), all of which are present at sites of inflammation, and involves the covalent attachment of heavy chains of IαI and PαI to the HMW HA molecules (Day and de la Motte, 2005; Zhuo et al., 2004). Free HA molecules fail to bind monocytes (de la Motte et al., 2003), suggesting that the binding of leukocytes to crosslinked HMW HA scaffolds occurs specifically through either induction of CD44 clustering or engagement of other co-receptors by the various molecules that adorn these HA cables. Importantly, the CD44-mediated leukocyte binding to HA scaffolds prevents their activation by controlling their ICAM-1-mediated interaction with the endothelium (Zhang et al., 2004; Selbi et al., 2004). A more recent study suggests that leukocyte interaction with HMW HA cables may actively induce growth factor and ECM secretion that promote tissue repair (Day and de la Motte, 2005). The crosslinked molecular network of HMW HA chains, typically found within joint tissues such as the articular cartilage surface of osteoarthritic knees, is also likely to prevent excessive loss of ECM and simultaneously guide the organization of new matrix (Milner and Day, 2003; Szanto et al., 2004). The crosslinked HMW HA scaffolds may also act as a reservoir for free radicals, thereby limiting excessive tissue damage (Zhuo et al., 2004; Rugg et al., 2005). This size-dependent effect of HA on tissue inflammation may explain, at least in part, the differences observed between fetal and adult tissue repair. Unlike adult wounds, early-gestation fetal wounds undergo scarless repair, which has been linked to the lack of an inflammatory response and consistently elevated levels of HMW HA resulting from increased synthesis by fetal fibroblasts (Chen et al., 1989) and decreased HAdase activity (West et al., 1997); exogenous addition of HAdase to such fetal wounds induces scar formation (Iocono et al., 1998). This, in accordance with the foregoing discussion, suggests that the HA fragments (LMW HA) in fetal wounds triggers inflammation through a TLR 4-mediated mechanism while the HMW HA likely exerts its conciliatory effect by attenuating inflammation via CD44mediated signaling.

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Interaction with tissue cells HA also exerts a strong biological effect by interacting with the CD44 and RHAMM (receptor for hyaluronan-mediated motility) receptors expressed on tissue cells. The binding of HA to these cellular receptors initiates downstream signaling that involves activation of protein phosphorylation cascades and cell cycle proteins, cytokine release and signal transduction to the cytoskeleton and nucleus that, together, regulate key cell functions such as adhesion, proliferation and migration (Entwistle et al., 1996; Toole, 2004). Importantly, tissue synthesis of HA is increased dramatically during important physiological processes such as morphogenesis and tissue repair. These physiological events involve proliferation and en masse movement of tissue cells, which is promoted by both HA-receptor interactions and the ability of HA to create large extracellular spaces required to accommodate huge cell population (Chen and Abatangelo, 1999; Toole, 1997). The expression of both HA and its receptors can be modulated by a variety of ECM signals, suggesting that the biological activity of HA is a tightly controlled phenomenon. In addition to facilitating cell migration and proliferation, HA also promotes matrix remodeling and prevents or minimizes wound scarring, likely due to combined cell signaling and physicochemical effects (Laurent et al., 1986a; Iocono et al., 1998). HA also interacts with CD44 and RHAMM receptors on endothelial cells and promotes angiogenesis, the process of new blood vessel formation. However, the overall effect depends on its molecular weight (MW); while the lower MW HA (oligosaccharide) promotes angiogenesis and new collagen deposition, higher MW HA inhibits new vessel formation (Dvorak et al., 1987; West and Kumar, 1989a; West and Kumar, 1989b; Lees et al., 1995). Although the exact mechanism for this MW regulation is not very clear, increased angiogenesis is accompanied by an increase in the levels of the HA-degrading enzyme, hyaluronidase (HAdase) (Liu et al., 1996; West and Kumar, 1989a).

28.3

Therapeutic implications of native hyaluronan

The excellent biological and physicochemical properties of native HA advocated its use in biomedical applications, although the early attempts exploited mostly its distinct viscoelastic behavior (Balazs and Denlinger, 1989; Weiss, 2000). One of the most common and successful applications of HA has been in the treatment of osteoarthritis, a pathological condition which is characterized by cartilage degeneration and subsequent loss of lubrication at the joints. Supplementation of exogenous HA to arthritic knees improves joint function and stability through increased retention of cartilage ECM molecules such as proteoglycans, which improves the viscoelastic

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properties of synovial fluid and suppresses cartilage degradation (Peyron, 1993; Ronchetti et al., 2000). The underlying mechanisms that mediate this therapeutic benefit of HA are not yet fully understood. Unmodified HA has also been used as an aid for ophthalmic surgery where its unique hygroscopic and viscoelastic properties help in creating large operative spaces and protecting the delicate corneal endothelium from physical damage (Laurent, 1981; Denlinger and Balazs, 1980). Although the initial therapeutic benefits were promising, it soon became clear that exogenous HA was not stable in the tissue for longer durations. This short residence time of unmodified HMW HA results from its spontaneous and rapid degradation by HAdase, which specifically cleaves the molecule at the β1,4 glycosidic bond (Stair-Nawy et al., 1999; Stern and Jedrzejas, 2006). HAdases are widely expressed in human tissues and degrade large HA chains to short oligos that are then metabolized by the surrounding tissue cells, which ensures proper turnover of tissue HA. Cartilage is a particularly dynamic tissue in terms of HA turnover, where the chondrocytes continuously synthesize and catabolize hyaluronan, with the typical half life of a hyaluronan molecules being ~2-3 weeks (Morales and Hascall, 1988; Ng et al., 1992; Flannery et al., 1998). Tissues that have access to lymph vessels, such as the skin and knee joint capsule, drain out excess HA through the lymphatic pathway, where the lined reticulo-endothelial cells actively eliminate the majority of HA, with the remainder catabolized by liver endothelial cells (Fraser et al., 1996; Fraser et al., 1997; Laurent et al., 1986b; Reed and Laurent, 1992); the half-life of HA in the bloodstream is only a few minutes (Fraser et al., 1984). The vigorous nature of HAdase activity is apparent from reports that estimate that almost one-third of the total hyaluronan in human tissues undergoes complete metabolic turnover during a single day. In addition to rapid degradation in vivo, unmodified HMW HA also lacks appropriate mechanical strength required to withstand mechanical loads that tissues such as cartilage commonly experience. This poor biomechanical property also makes handling HA very difficult.

28.4

Engineered hyaluronan

28.4.1 Chemical modification To improve biomechanical properties, increase residence time and allow easy handling, HA has been chemically modified and subsequently crosslinked to obtain a variety of stable derivatives. Importantly, although the HA derivatives exhibit improved physicochemical properties, they exhibit biocompatibility similar to the native HA; this makes HA a unique biomaterial for therapeutic applications since it combines the advantages of both naturally-occurring and synthetic materials (Allison and Grande-Allen, 2006; Vercruysse and

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Prestwich, 1998; Thierry et al., 2004). The carboxylic and hydroxyl groups on the HA backbone are the preferred targets for chemical modification. Esterification and carbodiimide-mediated reactions are the most common schemes that utilize the carboxylic group. The more commonly used HA esters, or HYAFF, are produced from the action of ethyl and benzyl alcohols on the carboxyl group, where the degree of hydrophilicity scales inversely with the degree of esterification (Campoccia et al., 1998). The carbodiimide reactions, on the other hand, involve activation of the carboxylic group at low pH (4.75) such that it couples efficiently to the multivalent dihydrazide groups such as adipic dihydrazide (ADH) and dithiobis propanoic dihydrazide (DTP), where the pendant hydrazide moieties can be further conjugated to other crosslinking or therapeutic agents (Prestwich et al., 1998; Vercruysse and Prestwich, 1998; Vercruysse et al., 1997). Typical derivatizations at the hydroxyl group include sulfation, esterification, and isourea coupling (Campoccia et al., 1998; Zhang and James, 2004; Abatangelo et al., 1997; Barbucci et al., 2000; Mlcochova et al., 2006). In addition to the carboxylicand hydroxyl-group modifications, HA can also be derivatized at its reducing end as well as the deacylated glucosamine groups, although these methods are not very popular due to low yield (Ruhela et al., 2006; Asayama et al., 1998; Dahl et al., 1988). Once HA is derivatized, it is crosslinked to produce robust biomaterials that exhibit the physical and degradation profiles desired for a specific biomedical application. Several crosslinking schemes have been developed that pair up with the appropriate derivatization method. They include, but are not limited to, diacrylate or divinylsulfone crosslinking, crosslinking via internal esterification, light, glutaraldehyde, carbodiimides and disulfides (Ghosh et al., 2006; Zheng Shu et al., 2004; Shu et al., 2002; Tomihata and Ikada, 1997; Park et al., 2003; Baier Leach et al., 2003; Crescenzi et al., 2003; Bakos et al., 2000; Campoccia et al., 1998; Sannino et al., 2004). Although intermolecular crosslinking is more common, intramolecular crosslinking is also possible, as seen in the autocross-linked hyaluronan (ACP™, Fidia), which is an ester derivative containing both inter- and intramolecular links between the hydroxyl and carboxyl groups (Mensitieri et al., 1996). Regardless of the approach, the covalent crosslinking of HA reduces its solubility in water such that addition of water causes the network to swell up to an equilibrium point where the osmotic swelling forces are balanced by the elastic forces of the internal atomic bonds. The strength and degradability of the HA derivative can be controlled by modulating both the degree of crosslinking and nature of the crosslinker and, therefore, it becomes possible to tailor these biomaterials for tissue-specific applications. For example, in osteoarthritic applications, the final product must be resilient and durable enough to withstand continuous cyclic mechanical forces, while biodegradation is a more crucial design parameter for cutaneous applications (Barbucci

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et al., 2002; Milas et al., 2001; Price et al., 2006). These HA-based biomaterials can subsequently be processed into myriad physical forms such as hydrogels, foams, films, fibers and microspheres, the final form depending on the nature of its eventual application (Figallo et al., 2007; Ji et al., 2006; Tang et al., 2007; Ghosh et al., 2006; Bae et al., 2006; Shu et al., 2002). The crosslinked HA network can also be coupled with various therapeutic agents that can be released at a local tissue site at a rate controlled by the degradation profile of the HA biomaterial (Lee et al., 2001; Yerushalmi et al., 1994; Wieland et al., 2007; Vercruysse and Prestwich, 1998).

28.4.2 Biological derivatization To promote tissue repair or regeneration, HA derivatives must activate tissue cells and induce them to migrate, proliferate and differentiate. Interestingly, although HA is known to interact with specific cellular receptors (CD44 and RHAMM), the HA-based biomaterials fail to support tissue cell adhesion and spreading, primarily due to the extreme hydrophilicity of HA that binds water layers on its surface and prevents protein deposition (Jackson et al., 2002; Sawada et al., 1999; Sawada et al., 2001). This issue has been addressed through biological derivatization of the chemically-modified HA. To do so, HA is first chemically modified such that multiple pendant groups are available for subsequent covalent coupling of biologically-derived cell-recognition peptides or proteins. The Arginine-Glycine-Aspartic acid (RGD) tripeptide sequence is the shortest biological motif used for HA modification since several cell-surface integrin receptors interact with the ECM via this peptide sequence. These HA-RGD hydrogels support extensive 3T3 fibroblast attachment, spreading and proliferation in vitro, and when these cells are encapsulated in the hydrogels and implanted in murine cutaneous wounds, they promote granulation tissue formation (Shu et al., 2004; Glass et al., 1996). Importantly, acellular HA-RGD hydrogels fail to support adult dermal fibroblast spreading and proliferation in vitro and fail to promote fibroblast invasion in vivo, suggesting that these hydrogels have good inductive but poor conductive properties (Ghosh et al., 2006; Shu et al., 2004). This limitation has been overcome by coupling more potent FN functional domains that simultaneously engage multiple cellular receptors (Ghosh et al., 2006). Compared to the HA-RGD hydrogels, these FN-modified hydrogels produced significant enhancement in primary adult human dermal fibroblast spreading, migration and proliferation, and more recent findings indicate that they also produce marked accentuation in porcine cutaneous wound repair. In addition to these FN-derived peptides, HA has also been modified by other polypeptides such as poly-L-lysine, poly-D-lysine, glycine or glutamine, and these derivatives show significant improvement in fibroblast adhesion and proliferation (Hu et al., 1999).

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To enhance bioactivity, HA has also been combined with larger adhesive proteins such as collagen and gelatin, with collagen receiving higher preference owing to its greater physiological significance and unique polymerization capability, which also improves the mechano-structural properties of the resulting blends. To improve stability in vivo, these blends are typically crosslinked; for example, composites have been developed by preparing HA-collagen coagulates, which are then crosslinked with starch dialdehyde and glyoxal to produce a collagenase-resistant and cell-interactive biomaterial (Rehakova et al., 1996). Other crosslinkers such as polyethylene oxide and hexamethylene diisocyanate can also be used (Soldan and Bakos, 1997). Similarly, gelatin has also been incorporated into HA solutions and the resultant mix crosslinked using carbodiimide (EDCI) chemistry; this blend subsequently promoted epidermal healing in vivo (Choi et al., 1999). In a more recent study, both HA and gelatin were identically derivatized using an EDCI chemistry that attached free pendant thiol groups to their backbones (Shu et al., 2003). When blended together, the thiol groups on HA and gelatin derivatives underwent spontaneous air-induced crosslinking to form stable, disulfide-linked composites that promoted extensive cell spreading and proliferation in vitro. Incidentally, although ECM-derived peptides are often required to promote cell adhesion and spreading on HA scaffold surfaces, no such biological derivatization seems to be necessary for three-dimensional (3D) cultures. For instance, when chick dorsal root ganglia were cultured in 3D hydrogels obtained by cross-linking thiol-derivatized HA, cultures produced robust neurite extension, which remained stable for up to eight days (Horn et al., 2007). A separate study showed that encapsulation of valvular interstitial cells (VICs), the most prevalent cell type in native heart valves, within crosslinked HA hydrogels maintained cell viability and promoted significant production of elastin over a period of six weeks (Masters et al., 2005). Furthermore, photo-crosslinked HA hydrogels have been shown to support chondrocyte viability, maintain the cells’ spherical shape, and promote extensive synthesis of cartilaginous matrix (Nettles et al., 2004). What regulates this difference in bioactivity between 2D and 3D cultures is not fully known although it may likely be due to enhanced intracellular signaling resulting from greater cell-HA interaction in a 3D environment. Just as various ECM-derived peptides or proteins are added to HA to improve its biological activity, HA is also added to a variety of polymeric materials to produce composites that retain the unique material properties of the synthetic polymer while exhibiting greater biological affinity. For example, HA-alginate composites can be formed in the presence of calcium, which facilitates gelation of alginate solution. The resulting gel exhibits both stable mechanical properties (due to alginate) and greater cell recognition (due to hyaluronan), which may together contribute towards increased ECM synthesis by encapsulated chondrocytes (Gerard et al., 2005; Oerther et al., 1999).

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Similarly, HA has also been blended with carboxymethylcellulose, an anionic polymer, and subsequently crosslinked to produce a robust biomaterial for prevention of postsurgical adhesions (Burns et al., 1996). In yet another variation, HA was added to hydroxyapatite-collagen mix that resulted in a biocompatible and mechanically robust material for use as a bone filler (Bakos et al., 1997).

28.4.3 Hemocompatibility In addition to being used as a scaffolding material for engineered tissues or an encapsulating material for drug delivery, HA is also preferred in applications that require continuous contact with blood. This is largely due to its ability to inhibit platelet adhesion and activation and delay both intrinsic and extrinsic coagulation pathways. These nonimmunogenic properties remain unaltered despite chemical modification of the HA network, which is often necessary to elicit greater response from vascular cells. For example, UV crosslinking of hylan or HA-divinyl sulfone (HA-DVS) gels renders them highly conducive to smooth muscle cell ingrowth without compromising their hemocompatibility (Amarnath et al., 2006; Ramamurthi and Vesely, 2005). The anticoagulant property of HA has been exploited in its use as a coating material for cardiovascular stents where it serves a dual role of: (a) increasing the stent’s hemocompatibility; and (b) releasing an encapsulated drug at a sustained rate. For instance, covalent immobilization of HA-heparin nanolayers on stainless steel cardiovascular stents not only improved the stent’s hemocompatibility but also promoted controlled release of a drug encapsulated within the HA-heparin complex (Huang and Yang, 2006). In another study, a HA-diethylenetriamine pentaacetic acid (DTPA) conjugate (HA-DTPA) was complexed with radionuclides yttrium and indium and used for coating stents and catheters during endovascular radiotherapy (Thierry et al., 2004). The resulting stents not only demonstrated significantly less fibrinogen adsorption and clotting but also maintained drug stability and release for over two weeks. It is important to note that although the chemical modifications of HA in these instances is performed solely to facilitate better binding to the stent, HA can also be derivatized such that it elicits differential anti-coagulant activity depending on the degree of modification. For example, Magnani et al. (1996) showed that an increase in the degree of sulfation of the hydroxyl group on HA disaccharide unit produces an increased resistance to the activation of factor Xa and thrombin, the components that trigger blood clotting cascade. Interestingly, the level of platelet aggregation follows an opposite trend, increasing with increasing degree of sulfation although even the highest aggregating effect is comparable with that of heparin (Barbucci et al., 1998).

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Implications for regenerative medicine

Stem cells, the primitive, multipotent cells that reside in the bone marrow and most adult tissues, undergo self-renewal and differentiation into multiple lineages, which together contribute to tissue homeostasis and repair (Weissman et al., 2001). These intrinsic properties of stem cells have led several groups to investigate their use in tissue engineering and regeneration applications. Notably, HA is being increasingly used as a scaffolding material for in vitro culture of stem cells prior to engraftment in the body. This is because: (a) in addition to being present in adult ECM, HA is also found in the bone marrow stroma (Wight et al., 1986) where it supports key functions of the resident mesenchymal stem cells (MSCs), including localization, proliferation and differentiation (Lee and Spicer, 2000); (b) HA is also found at high concentrations in the early embryonic ECM where it promotes gene expression and signaling, proliferation, migration and morphogenesis of embryonic stem cells (ESCs) (Toole, 2004); and (c) these stem cells express either one or both of the major HA receptors, viz. CD44 and CD168 (RHAMM) (Pilarski et al., 1999; Poulsom, 2007; Zhu et al., 2006). Therefore, HA is likely to elicit key stem cell functions necessary to maintain their therapeutic potential. Indeed, HA scaffolds have been shown to promote chondrogenic and osteogenic differentiation of MSCs both in vitro and in vivo when cultured in the presence of appropriate cytokines (Gao et al., 2001; Zavan et al., 2007; Facchini et al., 2006; Lisignoli et al., 2005; Kim et al., 2007). More detailed studies at the molecular level have shown that MSCs interact with HA scaffolds via a CD44-mediated mechanism and that this interaction causes differential expression of various chemokines and their receptors (e.g. upregulation of CXCR4, CXCL13 and MMP-3 while downregulation of CXCL12, CXCR5, MMP-13) that are involved in inflammation and matrix degradation (Lisignoli et al., 2006), the two processes that determine the outcome of tissue repair and regeneration. HA has also been shown to promote MSC adhesion and migration through a CD44-dependent pathway (Zhu et al., 2006), which has important implications in the design of regenerative strategies aimed at stimulating MSC homing to injury sites. Furthermore, HA hydrogels also provide a biocompatible environment for encapsulated human ESCs by maintaining the cells in an undifferentiated state while conserving their differentiation capacity under appropriate signals, as judged by their ability to form embryoid bodies in vitro (Gerecht et al., 2007). It is also likely that the bioactivity and compatibility of HA scaffolds can be further accentuated through systematic derivatization of the HA backbone, as is commonly performed for the more typical tissue engineering applications, as discussed earlier.

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Conclusion

HA is an important ECM component that plays a crucial role at various stages of a mammalian life span, from early embryogenesis to adult tissue hemostasis and repair. HA demonstrates unique viscoelastic properties, interacts with specific cell surface receptors that induce distinct intracellular signaling cascades, and modulates inflammation during tissue repair. In addition, HA can be easily modified to obtain more stable derivatives that are more resistant to enzymatic or hydrolytic degradation. Together, these properties have led to its widespread use in a variety of biomedical applications ranging from viscosupplementation to tissue engineering. Another striking feature of HA is its ability to resist activation of the blood clotting cascade, which renders it useful as a coating material for cardiovascular stents. That it influences both embryonic and mesenchymal stem cell function suggests that HA may likely play a vital role in regenerative medicine.

28.7

Future trends

Future success of HA-based therapies will depend on our ability to engineer smarter systems that address the complex, dynamic and reciprocal cell-ECM interactions that occur within the tissues. For example, LMW HA induces inflammation and angiogenesis that are essential for tissue repair while HMW HA inhibits both processes, which is important for controlling the reparative process. Therefore, if one were to deliver specially-derivatized LMW fragments that first initiate tissue repair and then, over time, form crosslinks to build HMW HA cables, it would be possible to modulate inflammation at a rate commensurate with tissue repair. Another approach could be to identify and utilize the cues from the wound to biologically and physiochemically modify HA such that following intravenous delivery at a remote site, it homes specifically to the injury site, builds a robust crosslinked scaffold and simultaneously activates the resident tissue cells to promote en masse cell ingrowth that is necessary for effective wound repair. To induce tissue regeneration, however, it would be desirable to develop HA derivatives that simultaneously activate and recruit stem and progenitor cells to the wound site. Recent reports that show that HA specifically interacts with stem cells via a CD44-mediated mechanism and promotes their migration and differentiation should serve as a platform for the development of such novel regenerative tools.

28.8

References

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Philipson L H, Westley J and Schwartz N B (1985), Effect of hyaluronidase treatment of intact cells on hyaluronate synthetase activity, Biochemistry, 24, 7899–906. Pilarski L M, Pruski E, Wizniak J, Paine D, Seeberger K, Mant M J, Brown C B and Belch A R (1999), Potential role for hyaluronan and the hyaluronan receptor RHAMM in mobilization and trafficking of hematopoietic progenitor cells, Blood, 93, 2918–27. Poulsom R (2007), CD44 and hyaluronan help mesenchymal stem cells move to a neighborhood in need of regeneration, Kidney Int, 72, 389–90. Prehm P (1983a), Synthesis of hyaluronate in differentiated teratocarcinoma cells. Characterization of the synthase, Biochem J, 211, 181–9. Prehm P (1983b), Synthesis of hyaluronate in differentiated teratocarcinoma cells. Mechanism of chain growth, Biochem J, 211, 191–8. Prestwich G D, Marecak D M, Marecek J F, Vercruysse K P and Ziebell M R (1998), Controlled chemical modification of hyaluronic acid: synthesis, applications, and biodegradation of hydrazide derivatives, J Control Release, 53, 93–103. Price R D, Das-Gupta V, Leigh I M and Navsaria H A (2006), A comparison of tissueengineered hyaluronic acid dermal matrices in a human wound model, Tissue Eng, 12, 2985–95. Ramamurthi A and Vesely I (2005), Evaluation of the matrix-synthesis potential of crosslinked hyaluronan gels for tissue engineering of aortic heart valves, Biomaterials, 26, 999–1010. Reed R K and Laurent U B (1992), Turnover of hyaluronan in the microcirculation, Am Rev Respir Dis, 146, S37–9. Reed R K, Lilja K and Laurent T C (1988), Hyaluronan in the rat with special reference to the skin, Acta Physiol Scand, 134, 405–11. Rehakova M, Bakos D, Vizarova K, Soldan M and Jurickova M (1996), Properties of collagen and hyaluronic acid composite materials and their modification by chemical crosslinking, J Biomed Mater Res, 30, 369–72. Ronchetti I P, Guerra D, Taparelli F, Zizzi F and Frizziero L (2000), Structural parameters of the human knee synovial membrane in osteoarthritis before and after hyaluronan treatment. In, Abatangelo G and Weigel P (eds), New Frontiers in Medical Sciences: Redefining Hyaluronan, Amsterdam, The Netherlands: Elsevier, 119–27. Rugg M S, Willis A C, Mukhopadhyay D, Hascall V C, Fries E, Fulop C, Milner C M and Day A J (2005), Characterization of complexes formed between TSG-6 and interalpha-inhibitor that act as intermediates in the covalent transfer of heavy chains onto hyaluronan, J Biol Chem, 280, 25674–86. Ruhela D, Riviere K and Szoka F C Jr (2006), Efficient synthesis of an aldehyde functionalized hyaluronic acid and its application in the preparation of hyaluronanlipid conjugates, Bioconjug Chem, 17, 1360–3. Sannino A, Madaghiele M, Conversano F, Mele G, Maffezzoli A, Netti P A, Ambrosio L and Nicolais L (2004), Cellulose derivative-hyaluronic acid-based microporous hydrogels cross-linked through divinyl sulfone (DVS), to modulate equilibrium sorption capacity and network stability, Biomacromolecules, 5, 92–6. Sawada T, Hasegawa K, Tsukada K and Kawakami S (1999), Adhesion preventive effect of hyaluronic acid after intraperitoneal surgery in mice, Hum Reprod, 14, 1470–2. Sawada T, Tsukada K, Hasegawa K, Ohashi Y, Udagawa Y and Gomel V (2001), Crosslinked hyaluronate hydrogel prevents adhesion formation and reformation in mouse uterine horn model, Hum Reprod, 16, 353–6. Scott J E and Heatley F (1999), Hyaluronan forms specific stable tertiary structures in aqueous solution: a 13C NMR study, Proc Natl Acad Sci USA, 96, 4850–5.

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Selbi W, De La Motte C, Hascall V and Phillips A (2004), BMP-7 modulates hyaluronanmediated proximal tubular cell-monocyte interaction, J Am Soc Nephrol, 15, 1199– 211. Shiedlin A, Bigelow R, Christopher W, Arbabi S, Yang L, Maier R V, Wainwright N, Childs A and Miller R J (2004), Evaluation of hyaluronan from different sources: Streptococcus zooepidemicus, rooster comb, bovine vitreous, and human umbilical cord, Biomacromolecules, 5, 2122–7. Shu X Z, Ghosh K, Liu Y, Palumbo F S, Luo Y, Clark R A and Prestwich G D (2004), Attachment and spreading of fibroblasts on an RGD peptide-modified injectable hyaluronan hydrogel, J Biomed Mater Res A, 68, 365–75. Shu X Z, Liu Y, Luo Y, Roberts M C and Prestwich G D (2002), Disulfide cross-linked hyaluronan hydrogels, Biomacromolecules, 3, 1304–11. Shu X Z, Liu Y, Palumbo F and Prestwich G D (2003), Disulfide-crosslinked hyaluronangelatin hydrogel films: a covalent mimic of the extracellular matrix for in vitro cell growth, Biomaterials, 24, 3825–34. Siegelman M H, Degrendele H C and Estess P (1999), Activation and interaction of CD44 and hyaluronan in immunological systems, J Leukoc Biol, 66, 315–21. Soldan M and Bakos D (1997), Complex matrix atelocollagen-hyaluronic acid. In Kukurová E, Advances in Medical Physics, Biophysics and Biomaterials, Stara Lesna, Slovak Republic, 58–61. Sorrell J M, Carrino D A, Baber M A and Caplan A I (1999), Versican in human fetal skin development, Anat Embryol (Berl), 199, 45–56. Stair-Nawy S, Csoka A B and Stern R (1999), Hyaluronidase expression in human skin fibroblasts, Biochem Biophys Res Commun, 266, 268–73. Stern R and Jedrzejas M J (2006), Hyaluronidases: their genomics, structures, and mechanisms of action, Chem Rev, 106, 818–39. Swann D A (1968), Studies on hyaluronic acid. I. The preparation and properties of rooster comb hyaluronic acid, Biochim Biophys Acta, 156, 17–30. Swann D A and Caulfield J B (1975), Studies on hyaluronic acid. V. Relationship between the protein content and viscosity of rooster comb dermis hyaluronic acid, Connect Tissue Res, 4, 31–9. Szanto S, Bardos T, Gal I, Glant T T and Mikecz K (2004), Enhanced neutrophil extravasation and rapid progression of proteoglycan-induced arthritis in TSG-6-knockout mice, Arthritis Rheum, 50, 3012–22. Tang S, Vickers S M, Hsu H P and Spector M (2007), Fabrication and characterization of porous hyaluronic acid-collagen composite scaffolds, J Biomed Mater Res A, 82, 323– 35. Taylor K R, Trowbridge J M, Rudisill J A, Termeer C C, Simon J C and Gallo R L (2004), Hyaluronan fragments stimulate endothelial recognition of injury through TLR4, J Biol Chem, 279, 17079–84. Teder P, Vandivier R W, Jiang D, Liang J, Cohn L, Pure E, Henson P M and Noble P W (2002), Resolution of lung inflammation by CD44, Science, 296, 155–8. Tengblad A, Laurent U B, Lilja K, Cahill R N, Engstrom-Laurent A, Fraser J R, Hansson H E and Laurent T C (1986), Concentration and relative molecular mass of hyaluronate in lymph and blood, Biochem J, 236, 521–5. Termeer C, Benedix F, Sleeman J, Fieber C, Voith U, Ahrens T, Miyake K, Freudenberg M, Galanos C and Simon J C (2002), Oligosaccharides of Hyaluronan activate dendritic cells via toll-like receptor 4, J Exp Med, 195, 99–111. Termeer C C, Hennies J, Voith U, Ahrens T, Weiss J M, Prehm P and Simon J C (2000),

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29 Biocompatibility of starch-based polymers A. P. M A R Q U E S, R. P. P I R R A C O and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

29.1

Introduction

Polymeric materials have a wide spread use in the biomedical field; they have been proposed to be used in Tissue Engineering (Cima et al., 1991), as delivery systems (Allen and Cullis, 2004; Colombo, 2000), suture materials (Echeverria and Jimenez, 1970; Herrmann et al., 1970), bone cements and screws (Bayne et al., 1975; Vert et al., 1985), as dental impression materials (Braden and Elliott, 1966) and in contact lenses (Poly, 1975) just to name some. However, one characteristic is common to all materials used in the biomedical field: biocompatibility. Biocompatibility of a material is generally defined as the ability of the material to perform within the host’s body without eliciting any immune response (Williams, 1992). In contrast to biocompatibility, other material’s characteristics depend on its intended application. For example, a bone Tissue Engineering (TE) scaffold must possess an adequate porosity and pore size, surface properties that promote cell adhesion, proliferation and the induction of neo-tissue formation, adequate mechanical and biodegradability properties (Hutmacher, 2000). However, in the case of drug delivery applications issues like the mechanical properties or bioinductivity (materials ability to induce new tissue formation) do not have the same importance (Allen and Cullis, 2004; Colombo, 2000). The properties of the chosen material for a given application will determine in great extent the properties of the developed device (Gomes and Reis, 2004; Vats et al., 2003). Among the materials considered promising for biomedical applications, the group of biodegradable polymers is one of the most studied due to their unique properties. This group of materials can be divided into two subgroups according to their origin: natural and synthetic. Each one of the sub-groups has advantages and disadvantages when their suitability for biomedical purposes is analyzed. In this brief introduction, general properties of natural-based vs synthetic biomaterials and their relevance on the biological performance of the materials and consequently on their biocompatibility will be focused. 738 © 2008, Woodhead Publishing Limited

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Synthetic biodegradable biomaterials such as poly(lactic acid) and poly(glycolic acid) were the first to have a widespread use in the biomedical field. Nowadays those materials remain the gold standard due to their regulatory approval which resulted from their use in other applications (Simamora and Chern, 2006). Nonetheless, other types of synthetic biomaterials, such as poly(ε-caprolactone), poly(propylene fumarate), poly(carbonates), poly(phosphazenes), and poly(anhydrides), since then have also been proposed for biomedical applications (Ben-Shabat et al., 2003; Choueka et al., 1996; He et al., 2001; Kweon et al., 2003; Laurencin et al., 1993; Mikos et al., 1994; Mooney et al., 1996). The main characteristic of synthetic biomaterials that renders them such a widespread use in the biomedical field is their processing and tailoring versatility (Gunatillake and Adhikari, 2003; Vats et al., 2003). In fact, depending on the material, a high degree of control of the material’s composition, structure and surface chemistry is possible to achieve, which allows an application-directed tailoring (Lai and Friends, 1997; Tiaw et al., 2007). However, issues regarding their performance in vivo, namely cytotoxicity issues and immunological issues deriving from their degradation products, directly arise from the synthetic nature of these materials (den Dunnen et al., 1997; Laaksovirta et al., 2002). As a consequence, an improving strategy encompasses the use of materials that mimic the physical, chemical and structural properties of living tissues. Such properties can be found in natural-origin biomaterials (Toole, 2004) obtained from either animal or vegetable sources. Among natural-origin biomaterials proposed for biomedical applications, collagen, (Murata et al., 1999; Ueda et al., 2002) fibrinogen (Haisch et al., 2000; Hojo et al., 2003), chitosan (Baran et al., 2004b; Zhang et al., 2000), hyaluronic acid (HA) (Liu et al., 1999; Solchaga et al., 1999), bacterial-derived poly(hydroxybutyrates) (Chen and Wang, 2002; Kostopoulos and Karring, 1994) and starch (Gomes et al., 2002; Malafaya et al., 2001; Santos et al., 2007) are the most common. One of the disadvantages attributed to these materials when compared to synthetic-origin biomaterials, concerns the difficulty in their processing and tailoring (Freier et al., 2005) and, in some cases, their low mechanical performance (Gomes et al., 2001b). The latter problem was addressed by researchers by blending the natural-origin materials with synthetic materials, thus reinforcing their mechanical properties (Lavik and Langer, 2004; Schmidt and Baier, 2000). Immunological issues that come from the fact that some natural-origin materials are contaminated with proteins derived from the source organism (Bruck, 1991; Piskin 1995), which can be resolved by the use of high purity grade materials, or due to their proteic nature, are also raised. The overall immunogenic potential of natural origin biomaterials is, nevertheless, very low and the advantages surpass the disadvantages. Most natural origin polymers are generally degraded in biological systems by hydrolysis, followed by oxidation, or enzymatically (Azevedo et al., 2003; Bruck, 1991; Rahmouni, 2001);

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contrarily, the majority of biodegradable synthetic polymers are not subjected to the action of enzymes and are hydrolysed by the action of water or serum (Bruck, 1991). Furthermore, in most cases the source of natural-based materials is almost unlimited, rendering researchers low cost availability of materials. However, the greatest advantage of natural-origin biomaterials is their range of properties that mimic some aspects of living tissues, allowing the host’s organism the possibility of recognition and metabolic processing, which is half way to achieving biocompatibility (Baran et al., 2004a; Espigares et al., 2002; Gomes et al., 2001a; Gomes et al., 2002; Malafaya et al., 2006; Silva et al., 2004a; Sousa et al., 2000). This resemblance to the living tissues is also reflected in the materials’ ability to induce the neo-formation of tissue which is, in general, higher than in synthetic materials. In this way, the increasing number of works that prove the high in vivo biocompatibility of several natural-origin based materials come as no surprise (Jeyanthi and Rao, 1990; Marques et al., 2005c; Salgado et al., 2007a).

29.2

Starch-based polymers in the biomedical field

Starch-based materials in particular, have been proposed for several applications in the biomedical field (Boesel et al., 2004; Espigares et al., 2002; Gomes et al., 2002; Oliveira et al., 2007; Salgado et al., 2007b; Silva et al., 2004a; Torres et al., 2007). Those applications are summarised in Table 29.1, highlighting the specific properties regarding each application that have been achieved and are considered critical for a specific cell response.

29.2.1 Bone-related applications In bone tissue engineering, starch has been used to produce scaffolds that serve as a template for transplanted cells to grow (Gomes et al., 2003; Salgado et al., 2004b) prior to subsequent implantation in vivo. These scaffolds must have general properties to be considered as a viable choice for TE purposes such as biocompatibility and biodegradability; however, they must also possess specific properties related to the characteristics of bone tissue. Bone tissue has two different macro-architectures. There are in fact two different types of bone tissue: spongy or trabecular bone and compact or cortical bone. Their proportion in the human skeleton is 20 and 80% respectively. Trabecular bone, where bone marrow is placed, is very porous (50–90% of porosity) while compact bone is much more dense than trabecular bone possessing only 10% of porosity, which reflects the fact that its modulus and ultimate compressive strength is around 20 times superior to that of trabecular bone. Moreover, bone is a hard tissue that results from a complex process of mineralization of the extracellular matrix performed by osteoblasts (Buckwalter et al., 1995). Thus, the development of a bone TE scaffold has

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Table 29.1 Relationship between key properties of starch-based materials and a specific biomedical application Applications

Assessed properties

References

Bone TE

Pore sizes ranging from 100–1000 µm Bone-related mechanical properties Support ECM production and mineralization

(Gomes et al., 2001; Gomes et al., 2003; Torres et al., 2007)

Vascularization potential

(Gomes et al., 2003; Mendes et al., 2003; Mendes et al., 2001; Salgado et al., 2007a; Salgado et al., 2005) (Santos et al., 2007)

Cartilage TE

Support Collagen Type II and GAGs production

(Oliveira et al., 2007)

Delivery Systems

Controlled degradation systems

(Baran et al., 2004; Coluccio et al., 2005; Devy et al., 2006; Malafaya et al., 2006; Silva et al., 2005b) (Devy et al., 2006; Echeverria et al., 2005; Malafaya et al., 2006; Silva et al., 2005a; Silva et al., 2007) (Echeverria et al., 2005; Silva et al., 2005a; Silva et al., 2007)

High loading capacity

Controlled release of the drug Spinal cord TE

Ability to support the proliferation of hippocampal neurons and glial cells

(Salgado et al., 2007b)

globally accepted requirements that have been pursued by researchers (Hutmacher, 2000; Lavik and Langer, 2004; Vats et al., 2003). An adequate porosity, with ideal pore size (200–900 µm) and interconnectivity to increase the surface to volume ratio, thus allowing cell in-growth and distribution throughout the porous structure, and promoting neovascularization (Gomes and Reis, 2004; Salgado et al., 2004a) is sought. Porosity is also determinant for a proper diffusion of nutrients and oxygen and for waste removal (Gomes et al., 2002; Gomes et al., 2003; Salgado et al., 2004b; Torres et al., 2007). Blends of starch with several other materials such as cellulose acetate (SCA), ethylene-vinyl alcohol (SEVA-C) and polycaprolactone (SPCL) processed by different methodologies have been studied (Gomes et al., 2001b; Gomes et al., 2003). The injection moulding-based method allowed the production of scaffolds comprising a porous core (with pore sizes ranging from 10 to 100 µm for SEVA-C-based materials and from 100 to 1000 µm for SCA) surrounded by a non-porous compact surface layer (Gomes et al., 2001b). SPCL scaffolds presenting very good porosity values, between 50–75%, with pore sizes ranging from 200–900 µm, thus were considered adequate for bone TE purposes (Gomes et al., 2003; Santos et al., 2007). Recently, Torres et al. (2007) also proposed several starch scaffolds produced by

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microwave processing, with pore size between 100 and 1000 µm and an indentation pressure, which correlates with the compressive strength, between 4–11 MPa. In fact, since bone is a tissue subjected to significant mechanical loads, the scaffold should have mechanical properties that match as much as possible those of bone and, at the same time, have a biodegradability rate that corresponds to the rate of new bone formation (Gomes et al., 2001b; Lavik and Langer, 2004). Another very important aspect is related to its surface. Surface chemistry and topography should be osteoconductive, promoting osteogenic cell adhesion, and osteoinductive, promoting osteoblastic progenitors and precursors recruitment to the implantation site (Salgado et al., 2005; Salgado et al., 2007a). The in vitro biological performance of starch-based scaffolds, assessed in a first stage with osteoblast-like cells (Gomes et al., 2001a; Salgado et al., 2002; Salgado et al., 2004b; Tuzlakoglu et al., 2005) and then with bone marrow cells (Gomes et al., 2003; Mendes et al., 2003; Tuzlakoglu et al., 2005) and human endothelial cells (Santos et al., 2007), revealed the great potential of these structures for bone tissue engineering applications. Osteoblastlike cells were able to adhere and proliferate onto the scaffolds filling some pores and to produce a skeletal structure typical of bone extracellular matrix as well as expressing typical osteoblastic markers, which is enhanced with the time of culture. Rat bone marrow cells were able to adhere and proliferate in SPCL fibre mesh scaffolds and, under a flow perfusion system, to be committed into the osteogenic phenotype, proving the osteoinductive potential of the scaffolds. Further studies (Santos et al., 2007) using the same fibre meshes and human micro-vascular and macro-vascular endothelial cells confirmed their adhesion and proliferation in the scaffolds while maintaining their phenotype, a requirement for the vascularization process. In vivo testing, using goat (Mendes et al., 2001) and rat (Salgado et al., 2007a) models, was also performed with some starch-based scaffolds. In the goat model, after implantation of SEVA-C scaffolds no significant immunological reaction was observed and satisfactory values of bone contact and bone remodelling were observed. In the case of the rat model, three different starch-based scaffolds, SEVA-C, SEVA-C/CaP and SCA, were implanted in critical size bone defects. All the scaffolds presented good integration with the host’s body, not eliciting a significant immunological response. Moreover some degree of new bone formation was seen in the scaffold/bone interface.

29.2.2 Cartilage tissue engineering The avascular character of cartilage tissue (Solchaga et al., 2001; Temenoff and Mikos 2000) reflects in some of its physiological properties. In fact, cartilage has a low metabolic profile and consequently very low selfregeneration ability. Articular cartilage in particular, acts as a weight bearing,

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shock absorbent tissue responsible for creating smooth gliding areas for the articulating skeleton to work properly. Its functionality depends to a great extent in maintaining the correct composition and structure of its highly hydrated ECM. The main components of the ECM of articular cartilage are collagen type II and proteoglycans, which are produced by the cells responsible for maintaining tissue homeostasis, the chondrocytes, although they only represent 1% of the total volume of cartilage tissue. These cells have a final developmental stage where they become hypertrophic chondrocytes and lose the ability to proliferate, becoming round and totally embedded in matrix. One of the problems in the regeneration of cartilage is the formation of fibrocartilage instead of a functional tissue. This type of tissue has diminished mechanical properties, compared to normal cartilage, resulting from deficient biochemical properties (Solchaga et al., 2001; Temenoff and Mikos, 2000). In cartilage tissue engineering the use of scaffolds is unavoidable. This strategy offers a 3D substrate for the anchorage-dependent chondrocytes to adhere to, since when cultured in a 2D environment these cells start to dedifferentiate (Temenoff and Mikos, 2000). In a comparative study (Oliveira et al., 2007), SPCL fibre mesh scaffolds were analysed in opposition to poly(glycolic acid) (PGA). Their suitability for cartilage TE purposes was assessed by in vitro culturing bovine chondrocytes under dynamic conditions, up to six weeks. Starch-based scaffolds presented a high degree of porosity (75%) and interconnectivity suitable for cell colonization throughout the structure of the scaffold and supported the production of Collagen type I and II as well as the production of glycosaminoglycans (GAGs).

29.2.3 Delivery systems In the case of many drugs, a constant blood concentration without significant fluctuations is critical for an optimal therapeutic performance. Drug delivery systems are structures that, due to their characteristics, can carry a drug and sustain its time- and space-controlled release (Pather, 1998). Therefore, the development of efficient drug delivery systems is of great importance. Various responsive polymers have been proposed as efficient carriers for drug delivery systems (Allen and Cullis, 2004; Colombo, 2000). There are two basic approaches for the use of polymeric matrices as delivery systems. One is based on the use of hydrophobic matrices that deliver an encapsulated drug as they erode or degrade by means of biological processes such as enzymatic degradation. The other is based on the use of hydrophilic matrices that can swell due to water uptake and deliver the drug through diffusion (Colombo 2000). Starch-based materials possess many favourable properties that allow their proposal as delivery systems. Their main feature concerns the natural degradation inside the human body by the action of amylolytic enzymes like

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alpha-amylase (Azevedo et al., 2003). Using native starch, this degradation is often too rapid in comparison to the drug delivery kinetics (Devy et al., 2006). Therefore, the blending of starch with synthetic polymers has been a suitable approach to avoid this issue. Starch-based delivery systems have been developed in several forms for controlled release applications (Baran et al., 2004a; Coluccio et al., 2005; Echeverria et al., 2005; Silva et al., 2007b). Coluccio et al. (2005) proposed the use of ethylene vinyl alcohol-starch-αamylase membranes as an enzymatic controlled drug delivery system. Also, Echeverria et al. (2005) investigated the potential of copolymers of ethylmethacrylate with starch (S-EMA) and with hydroxypropyl starch (HS-EMA), in the form of tablets, to serve as drug delivery systems. Starchbased hydrogels have also been studied as potential drug delivery systems by different groups (Baran et al., 2004a; Xu et al., 2006; Zhang et al., 2005). However, recent strategies have been focusing on the use of microspheres (Malafaya et al., 2006; Silva et al., 2004b; Silva et al., 2005a; Silva et al., 2007b). Their biocompatibility, shelf-life stability, high loading capacity, biodegradability, and controlled release of the encapsulated drug covers their major concerns and their use ranges from drug delivery to antigen delivery systems. Hydroxyethylstarch (HES) microparticles proposed as microcarriers for antigen delivery in immunotherapy approaches were tested in vivo, in a rat model, and shown to induce no immunological effects (Devy et al., 2006). Particles made of starch and poly-lactic acid (SPLA) and of SPLA and bioactive glass (SPLA/BG) were shown to support the adhesion and proliferation of MC3T3-E1 cells, without cytotoxicity (Silva et al., 2007a). Moreover, bioactive SPLA and SPLA/BG microparticles seemed to promote the osteogenic differentiation of the referred cells by themselves, a very positive property that could enhance the effect of a loaded osteogenic drug. In other work (Silva et al., 2005a), SPLA microspheres were loaded with Platelet Derived Growth Factor (PDGF), a known mitogen for several types of cells. Tests with MC3T3-E1 cells cultured in the presence of the microparticles demonstrated that there was a sustained release of PDGF into the culture medium that stimulated cell proliferation.

29.2.4 Spinal Cord Injury (SCI)-related applications The treatment of Central Nervous System (CNS) disorders is a challenging field mainly due to its low regenerative potential. Among these, Spinal Cord Injury (SCI) is one of the most frequent (Evans, 2001). The advances in tissue engineering have also provided new strategies to induce SCI regeneration, such as the development of a new generation of guidance scaffolds that promote axonal and nerve regeneration within its structure (Moore et al., 2006). Biodegradable materials from either synthetic or natural origin have been suggested as possible options for the development of SCI regeneration guidance

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tubes (Goraltchouk et al., 2005; Moore et al., 2006). Natural polymers such as collagen (Lin et al., 2006), dextran (Levesque et al., 2005) or chitosan (Freier et al., 2005) were already proposed for SCI related applications. A blend of starch with polycaprolactone (SPCL) has been recently proposed for this type of application (Salgado et al., 2007b). The viability and proliferation of central nervous system derived cells, such as hippocampal neurons and glial cells was tested in vitro in the presence of SPCL linear parallel filaments deposited on polystyrene coverslips. Both neuronal and glial cell populations were shown to adhere to the surface of SPCL filaments, without significant cell death and without major consequences on cell morphology and proliferation, which suggests the compatibility of the material surface properties over the tested cells. Moreover the SPCL based biomaterials did not seem to elicit an active response by the microglial cells, the inflammatory mediators of the central nervous system (CNS). In terms of functionality, despite the diminished responsive stimulus profile of the hippocampal neurons in comparison to the control samples, cells in the presence of the SPCL were found functional as the reduced response was attributed to the high cell densities obtained within the channels and consequently to the lower number of responsive cells, which delayed the establishment of the needed dendritic networks. Another interesting outcome of this study is related to the fact that oligodendrocytes, the myelinating cells of the CNS, were located in the vicinities of the SPCL filaments and not in the central areas of the channels, a property, according to the authors, that might be of use in order to stimulate myelination of newly regenerated nerves. Therefore, based on the absence of deleterious effects caused by the presence of the starch-based biomaterials on neuroprogenitor cells, suggesting that de novo neurogenesis is not compromised by their presence, it is possible to conclude that these materials are very promising for future studies on SCI regenerative medicine.

29.3

Cytocompatibility of starch-based polymers

Biocompatibility assessment covers several hierarchical stages each one of them aiming to evaluate the effect of different characteristics/properties of newly developed biomaterials, over the biological system. Besides guaranteeing that the role of the material is not compromised upon implantation, the emergence of novel biomaterials, in particular biodegradables, raises the question of eventual toxic effects of the metabolites resulting from the degradation process. The cytotoxicity screening of several starch-based blends has been mainly addressed using standardized cell lines, such as L929 (Gomes et al., 2001a; Marques et al., 2002; Mendes et al., 2001; Silva et al., 2004a), SaOs-2 (Marques et al., 2005a; Salgado et al., 2004b; Salgado et al., 2005; Salgado

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et al., 2002; Tuzlakoglu et al., 2005) and MC3T3 (Silva et al., 2007a). The biofunctionality of the starch-based polymers that are under consideration for use in orthopaedic temporary applications and as tissue engineering scaffolds has also been studied either with primary cultures (Oliveira et al., 2007; Salgado et al., 2007b; Santos et al., 2007) or bone marrow-derived cells (Gomes et al., 2003; Mendes et al., 2003; Silva et al., 2005b; Tuzlakoglu et al., 2005) obtained from different sources. Details on the performance of starch-based polymers for bone, cartilage and neuronal tissue regeneration were given in the previous section. Therefore, the cytotoxicity and the cytocompatibility of these materials, considering orthopaedic temporary applications, will be discussed in this section. The short-term effect of the degradation products of the SEVA-C blend and of SEVA-C/HA composites was initially evaluated (Mendes et al., 2001) and correlated with the presence of additives (ceramic fillers, blowing agents and coupling agents) and with the processing methods/conditions (Gomes et al., 2001a). In general, the obtained results showed that all the additives and the different processing methods required to obtain the different properties/ products, can be used without the inducement of cytotoxic behaviour by the developed biomaterials. The cytotoxicity of the extracts of other starchbased blends (SCA, SPCL and starch and SPLA70), as well as of the respective HA composites over different cell-lines, was evaluated also demonstrating promising results. (Marques et al., 2002; Marques et al., 2005) The worst performance of the blend of starch with cellulose acetate and HA composites was attributed to the high amount of low molecular weight chains and processing additives. The more severe thermal and shear cycles (extrusion compounding and injection moulding) that always provoke some thermal degradation (due to viscous heat dissipation) during the preparation of the composites, generates low molecular weight fragments (Reis et al., 1996) which are easily leached to the solution during the extract preparation. Nonetheless, these can be removed by an additional processing stage (Marques et al., 2002) leading to improved results. The biological performance of these starch-based materials was further assessed by establishing in vitro cell cultures in direct contact with the polymers and composites for different time periods in order to try to identify the effect of the surface of the materials over the normal cell metabolism. The adhesion, proliferation and morphology/spreading of osteoblast-like cells was shown to be influenced by the physico-chemical properties of the starch-based biomaterials (Marques et al., 2005a). Cells were well adhered and spread (Figure 29.1) on the majority of the surfaces showing slight differences in morphology that seem to disappear for longer culture times, which thus can be attributed to the differences in their surface properties. Due to their starch component, the materials in study have a high number of hydroxyl groups on their surfaces. It would be expected that the higher

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29.1 Scanning electron micrograph of osteoblast-like cells (SaOs-2) on the surface of SEVA-C polymer after seven days of culture.

hydrophilicity and oxygen content of SCA surface (Pashkuleva et al., 2005), would promote the adhesion of higher cell numbers. However, SEVA-C with the lowest oxygen content and a less hydrophilic (Pashkuleva et al., 2005) surface than SCA presented higher cell adhesion and a regular proliferation rate. The higher water uptake capability and the degradation rate of SCA seem to suggest that its surface experiences significant changes along the culture period which determine and influence cell behaviour. Moreover, cell proliferation rates were found to depend on the starch blend, thus on its synthetic component. In fact, the different percentages of starch and the miscibility of the starch-based blends might also have some influence in the biological performance of those biomaterials. SEVA and SCA, both with 50% of starch could be expected to induce a similar behaviour; however, the non-miscible character of the SCA blend can contribute to a completely different surface in terms of starch and synthetic component exposure and consequently cell adhesion. In addition the two starch blends with 30% of starch, SPCL and SPLA70, also presented very distinct cell adhesion results, which might indicate that in this case, the synthetic component rules cell adhesion and proliferation and that increasing the percentage of starch in the blend with polycaprolactone would improve those actions. In what concerns the biological performance of starch-based HA composites in comparison to unreinforced polymer, those materials induced more pronounced cell spreading, although different percentages of HA did not seem to significantly change osteoblast-like cell behaviour. The miscibility character of each one of the starch-based blends also determines the exposure of the HA particles within the samples.

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29.4

Natural-based polymers for biomedical applications

Immunocompatibility of starch-based polymers

The implantation of a biomaterial initiates a cascade of events, generally described as a foreign body reaction, which varies in time and in the inflammatory mediators involved (Anderson, 1988; Anderson, 1993). The duration and intensity of the response depends on several elements including the extent of the injury caused by the implantation procedure, factors related with the host (Colten, 1992; Emery and Salmon, 1991), and numerous properties of the implant such as chemical composition, surface free energy, surface charge, roughness, size and shape (Anderson, 1988; Malard et al., 1999; Marchant et al., 1990; Parker et al., 2002; Tang et al., 1998; Tengvall and Lundstrom, 1992; Yang et al., 2002). The emergence of biodegradable materials introduced more complexity to the biological response. Together with the foreign body reaction, the material is degrading, which may lead to changes in shape, surface roughness, release of degradation products (Gorna and Gogolewski, 2003; Hocker, 1998; Holland et al., 1990) and formation of particulates (Lam et al., 1995; Nakaoka et al., 1996); therefore, from the host perspective, potentially new elements to respond to. Many uncertainties are still present but some factors have been implicated in the occurrence and intensity of an inflammatory response against biodegradable implants. The difference in the rate of degradation and subsequently the difference in the kinetics of the release of the degradation products, such as monomers, oligomers and final fragments have been considered of major importance. The issue is that the velocity of degradation might be too fast allowing the inflammation process to take over, thus compromising the role of the device (den Dunnen et al., 1993; Gautier et al., 1998). The inflammatory cell reaction has been reported to be more intense for polymers that deteriorate rapidly (Fabre et al., 2001; Gibson et al., 1991; Winet and Bao, 1997). However, a too slow or hardly detectable degradation can also be undesired for some applications such as the use of biodegradables to support osteosynthesis. To date, a complete understanding of the biological responses to implanted biomaterials is still missing. The mechanisms of how a body reacts to implants over the course of time by inflammation, wound healing and the foreign body response are not fully understood. Immune system cells and chemical mediators are very important players in those reactions and are present at the implantation site, independently of the function of the device (Figure 29.2). Thus the evaluation of the mechanisms of inflammation, wound healing and foreign body reactions may provide useful information about the immunocompatibility of newly developed biomaterials. The factors that minimize inflammation will maximize biocompatibility. (Palumbo et al., 1997). The multiple responses possible during leukocyte activation and an

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incomplete understanding of their interactions, lead to the need to measure more than one response to characterize the extent of activation. A massive and generalized activation of leukocytes may impair the host by the excessive release of oxygen radicals, enzymes and/or cytokines (Figure 29.2). The activation of polymorphonuclear leukocytes (PMNs), the first cells to arrive at the implant site after surgery, may result in several processes such as chemotaxis, phagocytosis, degranulation and production of O 2– in a metabolic event known as respiratory burst (Chen et al., 1999; Nathan 1987). The degranulation of human PMNs after contact with starch-based materials and composites was assessed by means of quantifying the amount of lysozyme, a lysosomal protease, released into the culture medium (Marques et al., 2003). In that study, less than 20% of the potential maximum response was reached after incubation with the degradable materials. Moreover, lysozyme secretion was not significantly dependent on the material except for some SPCL composites. The effect of starch-based materials over PMNs activation was further analysed considering the oxygen-dependent mechanisms triggered upon cell activation (Marques et al., 2003). Changes in the free radical activity of the neutrophils were determined by measuring the luminescent response of Pholasin®, a photoprotein that emits light after excitation by reactive oxygen species. Two cell stimulants, formyl-methionyl-leucyl-phenylalanine (fMLP)

Biomaterial

IL-6 Blood vessels FGF PMN Monocytes/Macrophages IL-1 TNF-α

GM-CSF ROS

Lysozomal Enzymes

MHC II H2O2

O2

Fibroblasts

MHC II Dendritic cells

Lymphocytes

29.2 Schematic representation of the cellular and biochemical network involved in the reaction to foreign bodies such as a biomaterial

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and phorbol-myristate-acetate (PMA), with distinct mechanisms of action allowed conclusions about the respectively direct and indirect activation (via protein kinase C) of the NADPH oxidase system. The most striking result clearly demonstrated that the maximum response in the chemiluminescence tests was significantly reduced when the cells were exposed to the tested starch-based polymers when compared with the control. Four hypotheses were therefore defined in order to assess the nature of the phenomenon responsible for the reduction of the detected signal that was occurring at an early stage of the assay: (a) the event was a consequence of a cell/material interaction; (b) the free radicals were interacting with the material; (c) the material was quenching the light; on (d) the photoprotein Pholasin® was inhibited. The experiments set to demonstrate the hypotheses revealed that the observed reduction of the intensity of the signal was due to the polymeric nature of the materials which were capable of scavenging the free radicals, thus competing for the photoprotein. Furthermore, by changing the timeframe of the assay it was possible to observe the effect of the different adhesion behaviours induced by SEVA-C and SPCL surfaces also confirming the cell/material interaction effect upon PMN oxidative burst. Moreover, the intensity of the signal detected after fMLP stimulation was significantly lower than after PMA injection, indicating the activation of the NADPH oxidase both on the plasma membrane and on the secondary granules of the PMNs. The study of the immunogenic potential of starch-based polymers was also addressed, evaluating cell adhesion and differentiation of PMNs and a mixed population of monocytes/macrophages and lymphocytes (Marques et al., 2005b) after different culture times. This in vitro model was established in order to simulate aspects of the in vivo inflammatory response to evaluate individual and collective cellular effects resulting from the interaction of the different populations of inflammatory cells with the starch-based polymers. PMNs selectively adhered to the surface of the starch-based materials. Furthermore, their behaviour was also shown to be dependent on the time of culture and on the presence of ceramic. While SCA promoted higher PMN adhesion and lower activation, the number of cells from a mixed population of monocytes/macrophages and lymphocytes was found to be lower on that material, also showing a reduced amount of activated macrophages. In addition, the hydroxyapatite reinforcement induced changes in cell behaviour for some materials but not for others. However, HA generally showed reduced monocytes/macrophage adhesion and less potential to activate the cells. Probably a little unexpected, since it would be natural, is the maturation/ activation of monocytes/macrophages in an in vitro system where cells are exposed to foreign materials; but quite promising regarding the potential of starch-based materials, was the fact that the expression of ICAM-1 did not seem to be affected by the time of culture. The presence of HA down-

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regulated the maturation of monocytes into macrophages. Moreover, some composites down-regulated the expression of ICAM-1 molecules together with the expression of CD11b/CD18 integrins. In fact, the amount of activated macrophages (CD54 positive) was found to be lower in the presence of SCA. However, higher amounts of antigen-presenting cells (MHCII positive) were identified, at the time, in its surface which could be explained by the dynamics of the SCA surface resulting from the degradation process. The cellular activation potential of starch-based materials was further analysed considering that a chronic inflammatory response is mainly controlled, locally or systemically, by cytokines (Cohen and Cohen, 1996). The cytokine network is highly involved in attracting cells and inducing the production of cytokines as well as in guiding cellular functions. The release of interleukin1 beta (IL1-β), interleukin-6 (IL-6) and tumour necrosis factor-alpha (TNFα) after contact with starch-based materials was investigated as markers of early stages of injury/invasion. Moreover, the production of interferon-gamma (IFN-γ), recognized as a pro-inflammatory cytokine, was investigated. T lymphocyte mediated response was also addressed by quantifying interleukin4 (IL-4) and interleukin-2 (IL-2) (Marques et al., 2004). The results supported the hypothesis that different biodegradable polymers can affect mononuclear cell activation and the production of several cytokines associated with the inflammatory process. T-cells did not demonstrate significant activation. No IL-2 or IFN-γ was found in the culture supernatants contrarily to IL-6 which was detected in the highest amounts, for all the conditions, followed by TNF-α. IL-1β was produced in very low amounts, being undetectable in the presence of some of the starch-based materials. IL4 was the only cytokine that did not demonstrate any significant difference within the studied materials. The comparative analysis with a synthetic polymer showed that starch-based polymers (SCA presenting the most notorious result) and composites induced lower production of pro-inflammatory cytokines. Starch-based materials have been shown to be degraded by α-amylase (Azevedo et al., 2003) and phagocytosed by macrophages (Artursson et al., 1988; Desevaux et al., 2002b) inducing an excellent tissue reaction when implanted both in rats and mice (Desevaux et al., 2002a; Desevaux et al., 2002b). In works by other groups (Mendes et al., 2001; Souillac et al., 2001) starch-based materials implanted in rabbits and goats performed well and without adverse reactions. The host response to cross-linked high amylose starch (Contramid®) was found to be in accordance with the main phases of the inflammatory and foreign body responses to injuries caused by implanted devices (Anderson, 1988; Anderson, 1994; Ratner et al., 1996). After four months only a small residual scar was apparent macroscopically and it was even related with a less severe early reaction than a skin incision and closure with suture material sham (Desevaux et al., 2002b). The various in vitro models described above, and established to assess the

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reaction of immune system cells to the different starch-based blends (SEVAC, SCA, SPCL) and respective composites were validated in vivo using the rat subcutaneous model (Marques et al., 2005c). At implant retrieval no macroscopic signs of a considerable inflammatory reaction were observed and no cellular exudate was formed around the implants. A thin fibrous capsule, invariably containing inflammatory cells ranging from diffuse to concentrated density surrounded all implants. The types of cells recruited to the implantation site, as well as the specific subpopulations of activated cells were identified in order to try to understand the intensity of the tissue reaction. Recruited ED1 positive macrophages were present at the site of implantation and their number was increased for longer implantation times for some materials. Mature tissue macrophages (ED2) were only observed in the loose connective tissue surrounding the capsule of the implants and no significant differences were detected with time. In addition, although for some of the materials an abundant number of activated macrophages expressing ICAM1 could be identified, no foreign-body giant cells were present at the implantation site. Angiogenesis varied with the implantation times and also with the materials implanted. A significant increase in antigen-presenting phenotypes at the interface with some materials which can be associated with persistent local chronic inflammation was demonstrated. However, the almost complete lack of lymphocytes may be indicative of an innate mild foreign body reaction against these materials. SPCL and respective composites were the materials that stimulated the stronger tissue responses but generally biodegradable starch-based materials did not induce a severe reaction for the studied implantation times which contrasts with other types of degradable polymeric biomaterials, namely from synthetic origin. The in vivo observations validated the in vitro results confirming that the established in vitro models are reliable and can be used to estimate a potential inflammatory reaction provoked by newly developed biomaterials before implantation.

29.5

Conclusions

Starch is a natural-origin material that due to its promising properties has been proposed for a wide range of biomedical applications. Its degradable nature, associated with impressive mechanical properties, leads to the development of extensive work in the field of hard tissue replacement and regeneration. The potential of these materials in the biomedical area has been demonstrated by numerous published works that showed the different processing methodologies that can be applied to obtain starch-based biomaterials with distinct morphologies and mechanical and surface properties. In fact, the different starch-based materials have proven to perform well in vitro, and in certain cases in vivo, presenting a rather positive influence over

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a series of biological models. The outcomes of the biological tests performed with starch-based polymers showed that these materials do not induce a toxic effect over different cell types and also improve specific cell behaviour thus demonstrating the biofunctionality of the tested systems for each particular application. In what concerns the inflammatory/immune potential of the developed materials, the in vivo results validate the in vitro results confirming the suitability of the starch-based materials to be used in the biomedical field.

29.6

Acknowledgements

The authors would like to acknowledge to the European NoE EXPERTISSUES (NMP3-CT-2004-500283) and to the European STREP Project HIPPOCRATES (NMP3-CT2003-505758).

29.7

References

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hydrogen peroxide in response to products of macrophages and lymphocytes’, J Clin Invest, 80(6), 1550–1560. Oliveira J T, Crawford A, Mundy J M, Moreira A R, Gomes M E, Hatton P V and Reis R L (2007), ‘A cartilage tissue engineering approach combining starch-polycaprolactone fibre mesh scaffolds with bovine articular chondrocytes’, J Mater Sci Mater Med, 18(2), 295–302. Palumbo G, Avigliano L, Strukul G, Pinna F, Del Principe D, D’angelo I, AnnicchiaricoPetruzzelli M, Locardi B and Rosato N (1997), ‘Fibroblast growth and polymorphonuclear granulocyte activation in the presence of new biologically active sol-gel glass’, J Mater Sci Mater Med, 8(7), 417–421. Parker J A, Walboomers X F, Von den H J, Maltha J C and Jansen J A (2002), ‘Soft tissue response to microtextured silicone and poly-L-lactic acid implants: fibronectin precoating vs. radio-frequency glow discharge treatment’, Biomaterials, 23(17), 3545– 3553. Pashkuleva I, Marques A P, Vaz F and Reis R L (2005), ‘Surface modification of starch based blends using potassium permanganate-nitric acid system and its effect on the adhesion and proliferation of osteoblast-like cells’, J Mater Sci Mater Med, 16, 81– 92. Pather S I R, Syce J A and Neau S H (1998), ‘Sustained release theophylline tablets by direct compression Part 1: formulation and in vitro testing’, Int J Pharm, 164, 1–10. Piskin E (1995), ‘Biodegradable polymers as biomaterials’, J Biomater Sci Polym Ed, 6(9), 775–795. Poly I (1975), ‘Wettability of Hydrogels I. Poly (2-Hydroxyethyl Methacrylate)’, J Biomed Mater Res, 9, 315–326. Rahmouni M C, Nekka F, Lenaerts V and Leroux J (2001), ‘Enzymatic degradation of cross-linked high amylose starch tablets and its effect on in vitro release of sodium diclofenac’, Eur J Pharm Biopharm, 51, 191–198. Ratner B D, Hoffman A S, Schoen F J and Lemons J E (1996), Biomaterials Science. An Introduction to Materials in Medicine, San Diego: Academic Press. Reis R L, Cunha A M, Allan P S and Bevis M J (1996), ‘Improvement of the mechanical properties of hydroxyapatite reinforced starch based polymers through processing’, In Plastics in Medicine and Surgery-PIMS’96 (ed. J. Courtney), pp. 195–202. Institute of Materials, London, UK. Salgado A J, Coutinho O P and Reis R L (2004a), ‘Bone tissue engineering: State of the art and future trends’, Macromolecular Bioscience, 4(8), 743–765. Salgado A J, Coutinho O P and Reis R L (2004b), ‘Novel starch-based scaffolds for bone tissue engineering: cytotoxicity, cell culture, and protein expression’, Tissue Eng, 10(3–4), 465–474. Salgado A J, Coutinho O P, Reis R L and Davies J E (2007a), ‘In vivo response to starchbased scaffolds designed for bone tissue engineering applications’, J Biomed Mater Res A, 80(4), 983–989. Salgado A J, Figueiredo J E, Coutinho O P and Reis R L (2005), ‘Biological response to pre-mineralized starch based scaffolds for bone tissue engineering’, J Mater Sci Mater Med, 16(3), 267–275. Salgado A J, Gomes M E, Chou A, Coutinho O P, Reis R L and Hutmacher D W (2002), ‘Preliminary study on the adhesion and proliferation of human osteoblasts on starchbased scaffolds’, Mat Sci Eng C, 20, 27–33. Salgado A J, Sousa R A, Fraga J S, Pego J M, Silva B A, Malva J O, Neves N M, Reis R L and Sousa N (2007b), ‘Effects of starch/polycaprolactone based blends to be used

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30 Vascularization strategies in tissue engineering M. I. S A N T O S and R. L. R E I S, 3B’s Research Group, University of Minho, Portugal

30.1

Introduction

This chapter looks at vascularization in the tissue engineering field: the problems, state-of-the-art and strategies. Diffusion constraints in constructs and attempts to overcome this problem will be highlighted. We focus on bone regeneration, not only because it is one of the most complex tissues in which to build up a functional vascular supply but also due to the intimate relationship that exists between angiogenesis and bone formation/remodelling. Before describing current therapies to induce vascularization, the biological basis behind the structure, formation and maintenance of the vasculature will be discussed. Vascularization approaches will include stimulation of the endogenous angiogenic response (e.g. growth factors), pre-formation of structures that mimic the vascular tree (e.g. co-cultures), and both in combination. In conclusion, we will consider the evaluation of angiogenesis in tissue engineered products and present two of the most common in vivo models.

30.2

Biology of vascular networks – angiogenesis versus vasculogenesis

New blood vessel formation can occur through two distinct processes: angiogenesis and vasculogenesis. While angiogenesis is the formation of blood vessels from existing ones, vasculogenesis is the de novo formation of blood vessels (Patel and Mikos, 2004). Vasculogenesis is responsible for the formation of the capillary structure of the primary vascular plexus during embryonic development, although it has been noted in adults as well (Cassell et al., 2002). The few physiological situations in the adult where microvascular remodelling is required include the female menstrual cycle and skeletal muscle responding to exercise (Ferrara and Alitalo, 1999). Angiogenesis is related to pathological situations, such as inflammation, wound healing, ischemia and hypoxia (Kannan et al., 2005). In addition, 761 © 2008, Woodhead Publishing Limited

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many diseases are driven by persistent unregulated angiogenesis, such as arthritis, where new capillary blood vessels invade the joint and destroy cartilage; diabetes, where new capillaries in the retina invade the vitreous humour, bleed, and cause blindness; and tumour growth, which is dependent on a vascular bed to continue to proliferate (Folkman and Shing, 1992). Angiogenesis in vivo is regulated by complex interactions between endothelial cells (ECs) and other cell types (e.g. monocytes/macrophages, fibroblasts, smooth muscle cells/pericytes, osteoblasts) and other compounds such as cytokines and growth factors, molecules of the ECM, and cell surface adhesion molecules (Peters et al., 2002). The sprouting process consists of several consecutive steps: (1) local degradation of the basement membrane on the side of the venule closest to the angiogenic stimulus; (2) migration of ECs toward the angiogenic stimulus; (3) alignment of ECs in bipolar mode; and (4) formation of a lumen (Patan, 2000). This is a complex process mediated by a multitude of growth factors, for instance, angiogenic sprouting is mediated by transforming growth factor-β (TGF-β), while maturation of vessels is via angiopoietin-1 and -2 (Ang-tie) and platelet-derived growth factor (PDGF) pathways (Kannan et al., 2005).

30.3

Vascularization: The hurdle of tissue engineering

Since it was first defined in 1993 (Langer and Vacanti, 1993), the field of tissue engineering has evolved in an exponential way. Nevertheless, and despite all the great achievements, tissue engineering products are mainly limited to regeneration of avascular and tissues with low metabolic demands, such as cartilage, or tissues with small two-dimensional volumes, such as skin (Lokmic et al., 2007). Successes have also been reported for the regeneration of the neo-bladder and heart valve, but this is in part due to the fact that these structures are thin enough to survive on diffusion nutrition until a blood supply is established (Cassell et al., 2002; Atala et al., 2006). In metabolically active tissues, such as trabecular bone, bone marrow and liver, the distance that oxygen must diffuse between a capillary lumen and a cell membrane is usually 40 to 200 µm (Muschler et al., 2004; Lee et al., 2006). To date, the majority of implants have relied solely upon diffusion initially, and in the later stages on post-implantation vascularization (Kannan et al., 2005). When the vessels that deliver oxygen are initially confined to the outer surface of the graft, the metabolic demands of transplanted cells, particularly those deeper in the scaffold, are not met and the successful integration of the implant is jeopardized (Muschler et al., 2004) (Fig. 30.1). Of all the metabolites, oxygen is the limiting factor in cell survival in most grafts, due to its low transport coefficient through the aqueous environment of living tissues and high consumption (Muschler et al., 2004; Lee et al.,

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Osteoblasts Apoptotic osteoblasts

30.1 Scheme illustrating diffusion constraints of a thick scaffold for bone regeneration.

2006). Coger’s group has described a technique for enhancing O2 transport through aqueous extracellular matrix (ECM) gels aimed at liver reconstruction. On normal collagen gels O2 transport is diffusion-dominant, thus restricting hepatocyte viability and function. An enhanced collagen-type ECM was developed by the addition of microporous beads in which O2 is mutually transported by diffusive and convective flows, consequently inducing higher hepatocyte function levels (McClelland and Coger, 2000; McClelland et al., 2003; Lee et al., 2006). The mass transport in a graft, defined as the in and out movement of substrate molecules and products of metabolism, is highly impaired by the thickness of the implant. For instance, a 1 cm thick scaffold can support 280 000 cells/cm3 without central necrosis, whereas in native autogenous cancellous bone this value is 1000-fold higher (~5 × 108 cells/cm3) (Muschler et al., 2004). The key difference is an established vascular network in bone that supplies cells with all the required nutrients. Capillaries provide an effective means of mass transfer because their small diameter, approximately 6–8 µm, ensures that the residence time of the blood is greater than or equal to the radial diffusion time of the chemical species within the tissue (Freed and Vunjak-Novakovic, 1998). Therefore, the size of a scaffolding material without a functional vascular network is limited by mass transfer constraints (Freed and Vunjak-Novakovic, 1998). Hence the successful application of tissue engineering therapies depends on the development of new strategies that augment vascularization.

30.4

Neovascularization of engineered bone

Successful vascularization leading to tissue regeneration can only be achieved based on a deep understanding of formation, regeneration, the growth factors and cytokines that orchestrate the processes, and the role that angiogenesis plays in the overall process. Therefore, before going into detail about strategies to establish a functional vascular supply, here we discuss the relevance of angiogenesis in osseous formation and repair. Bone is a vascularized tissue composed of an organic matrix consisting of type I collagen and other proteins, and inorganic mineral consisting mainly of carbonate-rich hydroxyapatite (Freed and Vunjak-Novakovic, 1998). It contains a variety of different cell types: vascular cells, marrow cells, pre-

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osteoblasts, osteocytes, chondroblasts and osteoclasts, all executing distinct cellular functions to allow the bone to work as a highly dynamic organ (Meyer et al., 2004). A complex network of blood vessels, intraosseous circulation, assures the metabolic survival of these cells, allows traffic of minerals between the blood and bone tissue and sends the blood produced within the bone marrow into the systemic circulation (Laroche, 2002). The blood supply of long bones is provided by several groups of arteries: proximal/ distal metaphyseal arteries, proximal/distal epiphyseal arteries, diaphyseal nutrient arteries and periosteal arteries (Carano and Filvaroff, 2003). Nevertheless, angiogenesis does not just play a passive role in providing substrates for the process of osteogenesis, but it precedes osteogenesis in many practical situations – blood vessels play an active role in the process of osteogenesis (McCarthy, 2006). Regarding osseous formation, there are two distinct processes (Gerber and Ferrara, 2000). The first is intramembranous bone formation, during which mesenchymal cells condense and directly differentiate into osteoblasts to deposit bone matrix. The second is endochondral bone formation, during which a cartilage mould is first formed from mesenchymal condensations, and is then replaced by bone and bone marrow (Chung et al., 2004). For instance in intramembrenous bone formation, extensive vascularization is observed at the transition of pre-osteoblasts to osteoblasts (Deckers et al., 2002). In endochondral bone formation, an avascular cartilage template is replaced by highly vascularized bone tissue (Maes et al., 2002; Gerber and Ferrara, 2000). The progression of endochondral bone formation is dependent on efficient angiogenesis, and is blocked if angiogenesis is blocked, as illustrated by both experimental and pathological conditions (Bianco et al., 2001). Thus it is not surprising that cytokines and growth factors that regulate intraosseous angiogenesis also regulate bone remodelling, and close links exist between blood supply and bone formation and resorption. For instance, most diseases characterized by increased bone resorption are associated with increased bone vascularization (Laroche, 2002). As regards repair of fractures by callus production, four overlapping phases have been identified. In the first phase following injury, disruption of blood vessels leads to the formation of a haematoma. Then, ECs migrate from preexisting blood vessels in a directional manner towards a chemotactic stimulus and form the soft callus (Probst and Spiegel, 1997). Then the callus becomes mineralized, creating hard callus. Finally, the large fracture callus is replaced with secondary lamellar bone and the vascular supply returns to normal (Carano and Filvaroff, 2003). It is unfair to regard one component of a multicomponent biological system to be more or less important than another; however it is necessary to emphasize the significance of blood vessel formation for bone formation and repair. Accordingly, strategies that enhance vascularization or angiogenesis should benefit bone wound repair (Orban et al., 2002). © 2008, Woodhead Publishing Limited

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30.5

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Strategies to enhance vascularization in engineered grafts

30.5.1 Mature and progenitor endothelial cells ECs form the inner lining of a blood vessel and provide an anticoagulant barrier between the vessel wall and blood. In addition to their role as a selective permeability barrier, ECs are unique multifunctional cells with critical basal and inducible metabolic and synthetic functions (Sumpio et al., 2002). They are also the main cell type involved in angiogenesis, and are therefore the main cell target for vascularization strategies. The first EC culture was established in the early 1970s (Jaffe et al., 1973). ECs are isolated from freshly obtained human umbilical cords by collagen digestion of the interior of the umbilical veins (Jaffe et al., 1973) (Fig. 30.2a). Umbilical veins are probably the most widely used source for human ECS, since they are more easily available than many other vessels, they are free from any pathological processes and they are physiologically more relevant than many established cell lines (Marin et al., 2001). In fact, most of our knowledge of ECs comes from the study of human umbilical vein endothelium, but there are some drawbacks associated with cells derived from macrovasculature. Once they are close to senescence, the cells’ response to growth factor becomes weaker due to defective signalling pathways (Garlanda and Dejana, 1997). Therefore, in approaches involving growth factors, it is better to choose cells derived from microvasculature, the most common sources of these cells being skin, fat tissue and juvenile foreskin. Human microvascular ECs can be isolated from various human tissues, including foreskin, adult dermis, lung, glomerulus, endometrium, and brain (Laurens, 2004) (Fig. 30.2b). Independent of the approach adopted to initiate vascularization, the success of the implant depends on an appropriate response from the host vasculature. Therefore, endothelialization of scaffolds, i.e. culture in monolayer with ECs, is important insofar as it indicates if the scaffolding material has elicited an adequate response from ECs. For instance, Santos et al. (2007) examined the ability of fibre meshes made from a blend of starch with polycaprolactone (SPCL, 30/70% wt), a scaffold for bone repair, to serve as an appropriate substrate for ECs, using HUVECs as representative of macrovasculature and the microvascular cell line HPMEC-ST1. The in vitro results showed not only the expression of constitutive markers, such as PECAM, VE-cadherin and vWF, but also of inducible markers related to the inflammatory response in the presence of the right stimulus (Fig. 30.3). Table 30.1 sums up the most common endothelial markers and their respective function. The in vitro culture with ECs can also be used to pre-form a capillary-like structure to significantly accelerate neovascularization of the graft. A hyaluronan-based biomaterial for skin regeneration was cultured with HUVECs and the cells reorganized into a microcapillary network inside the dermal substitute (Tonello et al., 2003). © 2008, Woodhead Publishing Limited

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(a)

(b)

30.2 (a) Monolayer of human ECs derived from umbilical cord (HUVECs); (b) Human dermal microvascular endothelical cells (HDMECs) isolated from juvenile foreskin (x10).

A unique characteristic of ECs is that, although they present many common functional and morphological features, they also display remarkable heterogeneity in different organs (Garlanda and Dejana, 1997). Chi et al. explored EC specialization on a global scale, using DNA microarrays to determine the EC expression profile. They found that ECs from different blood vessels and microvascular ECs from different tissues have distinct and characteristic gene expression profiles (Chi et al., 2003). Therefore, it is important to emphasize the phenotypic variation between ECs in different portions of the vascular tree, and between arterial and venous cells, such that

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300 µm

30.3 Confocal laser scanning microscopy of SPCL fiber-mesh scaffold seeded with HDMEC and stained for PECAM-1. Nuclei were counterstained with Hoechst.

cells from different locations within an individual not only express different markers but can also generate different responses to the same stimulus (Sumpio et al., 2002). Bone marrow, fat tissue and peripheral blood in adults contain a special sub-type of progenitor cells which are able to differentiate into mature ECs, thus contributing to re-endothelialization and neovascularization. These angiogenic cells have the properties of embryonic angioblasts and are termed endothelial progenitor cells (EPCs) (Hristov et al., 2003). These precursors are identified through the expression of three cell markers (CD133, CD34, and the VEGF-R2), and have the capacity to proliferate, migrate and differentiate into ECs, but have not yet acquired characteristic mature endothelial markers (Luttun et al., 2002; Hristov et al., 2003). EPCs can be grown from purified populations of CD34+ or CD133+ haematopoietic cells, purified CD14+ monocytes, or total peripheral blood mononuclear cells (Urbich, 2004). There is a population of EPCs in peripheral blood, the so-called outgrowth endothelial cells (OECs), which have a cobblestone-like morphology suggestive

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Table 30.1 Constitutive and inducible endothelial markers Classification

Cell adhesion molecules (constitutive expression)

Marker

Function

Reference

Bon Willebrand factor

Mediates adhesion and aggregation of platelets at sites of vascular injury

(Ruggeri, 2000)

PECAM

Occurs on EC membrane close to the intercellular junctions and regulates the adhesion of ECs to other cells of the same type and to leukocytes Mediates cell-cell contact between ECs and plays a relevant role in the maintenance of vascular integrity

(Cenni et al., 1997)

VE-cadherin

Cell adhesion molecules (inducible expression related with inflammatory response)

E-selectine

ICAM

VCAM

Induces a prolonged contact between circulating leukocytes, resulting in a decelerated rolling along the endothelium During inflammation is up-regulated several fold to facilitate EC-leukocyte adhesion, especially neutrophils and monocytes Favors the adhesion and transendothelial migration especially of lymphocytes

(Nachtigal et al., 2001)

(Muller et al., 2002)

(Remy et al., 1999)

(Cenni et al., 1997)

of the endothelial phenotype and express several endothelial markers. These cells have a high proliferative capacity, expansion in optimal numbers and phenotypic stability in long-term cultures, making them an optimal candidate for autologous cell therapies (Fuchs et al., 2006a). OECS have been used in combination with fibroin silk fibre meshes for applications in tissue engineering. The results showed endothelialization of fibroin silk fibre meshes, maintaining their endothelial characteristics and functions (Fuchs et al., 2006b).

30.5.2 Matrices Biomaterials play a critical role in the engineering of new functional tissues for the replacement of lost or malfunctioning tissues. They provide a temporary scaffolding to guide new tissue growth and organization (Kim et al., 2000). Both natural and synthetic biomaterials have been explored as matrices or scaffolds for therapeutic angiogenesis (Zhang and Suggs, 2007). Table 30.2 lists several natural scaffolding materials used in diverse vascularization strategies for tissue engineering, including incorporation of growth factors,

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Table 30.2 Examples of natural biomaterials used in tissue engineering strategies to increment and accelerate the establishment of the vascular network Natural biomaterial

Tissue to regenerate

Strategies

References

Silk fibroin

Several, mainly soft tissues

Endothelialization, co-culture

(Unger et al., 2004b, Unger et al., 2007, Unger et al., 2004a, Fuchs et al., 2006b)

Hyaluronic acid

Dermis

Endothelialization

(Tonello et al., 2003)

Fibrin

Diverse

Growth factors delivery

(Ehrbar et al., 2004)

Chitosan

Bone

Growth factors delivery

(Elcin et al., 1996, Lee et al., 2002)

Collagen

Heart or diabetic ulcers

Growth factors delivery, co-culture

(Nillesen et al., 2007, Koike et al., 2004)

Gellatine

Soft tissues

Endothelialization

(Dubruel et al., 2007)

Starch-based

Bone

Endothelialization

(Santos et al., 2007)

endothelialization and co-culture. Among natural biomaterials, fibrin, collagen and gelatine are the most widely investigated. The choice of these materials relies mainly in the fact that they are natural elements occurring in the microenvironment of ECs in vivo. But to design adequate scaffolds, besides selecting the appropriate materials and routes to process them, it is also necessary to consider the respective porosity, interconnectivity and surface characteristics (Gomes and Reis, 2004). An ideal scaffold architecture would include an artificial capillary network that could anastomose with the patient’s own blood vessels during surgery (Griffith and Naughton, 2002). This capillary bed would include small arteries (1–2 mm) conducting blood into an arteriolar network (100–1000 µm), which would eventually end in capillary-like vessels of 10–15 µm. The end capillaries would not be more than 100–200 µm from every cell, and should convert into a venous collecting system for venous blood (Kannan et al., 2005). Some advances have already been made in this direction. Borenstein et al. (2002) produced organ templates with feature resolution of 1 micron, well in excess of that necessary to fashion the capillaries necessary for microcirculation in the organ. They used advanced microfabrication technologies, such as silicon micromachined template wafers. Following the same concept of engineering a vasculature using microfabrication in silicon, Shin et al. (2004) developed a methodology to create an endothelialized network with a vascular confined geometry. Despite the great potential of microfabricated scaffolds, the technology is still very new and so far only synthetic materials have been used.

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Another architecture for bone tissue engineering, a nano/micro fiber combined scaffold, has been proposed by Tuzlakoglu et al. (2005). The innovative structure of the nano/micro fiber scaffold is inspired by ECM, which combines a nano-network, aimed at promoting cell adhesion, with a microfibre mesh that provides the mechanical support. In vitro studies with ECs have proven the capacity of this structure to elicit and guide the threedimensional distribution of ECs (Santos et al., 2007) (Fig. 30.4). Vaz et al. (2005) described a scaffold architecture mimicking that of a blood vessel morphologically and mechanically by sequential multilayering electrospining. The innovation and potential of this technique is to design scaffolds with a hierarchical organization through a layer-by-layer process and maintain control over fibre orientation. Porosity is another important factor to consider in bone regeneration. Ripamonti et al. (1992) demonstrated that pore sizes of 150 µm do not support vascularization. For bone regeneration, it is recommended that pores sizes are >300 µm, due to enhanced bone formation and capillaries. Because of vascularization, pore size has been shown to affect the progression of osteogenesis. Small pores favoured hypoxic conditions and induced osteochondral formation before osteogenesis, while large pores which are well vascularized lead to direct osteogenesis (without preceding cartilage formation) (Karageorgiou and Kaplan, 2005).

30.4 Scanning electron microscopy of nano/micro fiber combined scaffold with HUVEC cells after seven days of culture.

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30.5.3 Incorporation of angiogenic growth factors Several natural polymers have been used as delivery matrices for angiogenic growth factors, such as alginate (Tanihara et al., 2001), chitosan (Elcin et al., 1996), fibrin (Sakiyama-Elbert and Hubbell, 2000) and collagen (Pieper et al., 2002; Lee et al., 2002). Besides the immense therapeutic potential of this approach, several aspects have to be taken into consideration. One of them is the complexity of blood vessel formation. Angiogenesis is a complex process mediated by a multitude of growth factors, therefore tissue regeneration cannot rely on the delivery of single factors as typically happens (Richardson et al., 2001) (Table 30.3). In addition, there are hazards associated with treatment with angiogenic growth factors, including concerns that pathologic processes dependent on angiogenesis, for instance tumour development, atherosclerosis, and proliferative retinophathies, may be exacerbated (Moldovan and Ferrari, 2002). Thus, an effective and safe angiogenic therapy is not only dependent on the right combination of growth factors delivered and/or administered, but also on ensuring that release is controlled in a temporal and dose manner. Of all the angiogenic growth factors, VEGF is probably the most essential for the development and differentiation of the vascular system and therefore the one that has been most extensively studied in drug delivery systems (Ferrara and Alitalo, 1999). In spite of this, its application in therapy remains difficult because blood vessels formed by exposure to high doses of VEGF tend to be malformed and leaky (Zisch et al., 2003). Ehrbar et al. (2004) Table 30.3 Main angiogenic growth factors Angiogenic growth factor

Action

References

Angiopoietin 1

Important role in the assembly of newly formed vasculature and in the maintenance of vascular integrity

(Yamakawa et al., 2003)

Angiopoietin 2

Is a natural antagonist of Ang1. Can cause endothelial cell apoptosis and vascular regression, but in the presence of VEGF destabilizes the preexisting vasculature making it more responsive to angiogenic stimuli

(Yamakawa et al., 2003)

VEGF

Causes blood vessel hyperpermeability and acts specifically on endothelial cells to induce their migration and proliferation.

(Ferrara et al., 2003)

FGF

Initiate the basement degradation cascade, stimulate EC migration and migration. Nevertheless, FGFs are not specific for ECs and act on a variety of cell types

(Patel and Mikos, 2004)

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avoid this problem by loading a fibrin scaffold with an engineered variant form of VEGF, alpha2PI1-8-VEGF121. As happens in nature, this engineered growth factor has the ability to bind to ECM components, keeping them immobilized until released by local cellular enzymatic activity. Angiogenesis is a multi-step process and there are different growth factors involved in the different stages; therefore more complex approaches with multiple growth factors are closer to the in vivo situation. Some systems for dual growth factor delivery have been described. For instance, Nillesen et al. (2007) described a strategy in which acellular collagen scaffolds were loaded with FGF2 and VEGF. The scaffolds were implanted subcutaneously in rats and the results indicated that addition of both angiogenic growth factors enhanced early mature vasculature in relation to the non-loaded acellular collagen. Richardson et al. (2001) reported a new polymeric system that allows for the tissue-specific delivery of two or more growth factors, with controlled dose and rate of delivery. Another major drawback of therapies based on peptide growth factors, besides despoliation of pathologic processes, is their high cost and susceptibility to aggregation and degradation. One possible solution might be the use of synthetic molecules such as phthalimide (PNF1) (Wieghaus et al., 2006). This was the first synthetic small-molecule inducer of angiogenesis reported. A body of knowledge is still being built around this molecule, but it has already been proposed that the pro-angiogenic mechanism of PNF1 is associated with TGF-β signalling pathways (Wieghaus et al., 2007).

30.5.4 Co-culture systems Scaling up the bone tissue vascular supply is of major importance for clinical applications. One approach to overcome this serious and currently still not definitively solved problem is the development of co-cultures of bone cells and vascular cells. Osteogenesis has long been associated with angiogenesis; in fact since the 18th century, when Hunter reported for the first time that blood vessels are key contributors to the process of osteogenesis, both in development and during repair (Carano and Filvaroff, 2003). Bone development and remodelling depend on complex interactions between bone-forming osteoblasts and other cells present within the bone microenvironment, particularly ECs, that may be pivotal members of a complex interactive communication network in bone (Guillotin et al., 2004; Choi et al., 2002). In a pioneering paper, Levenberg et al. (2005) have shown that it is indeed possible to vascularize an engineered construct. They demonstrated the potential of such an approach by engineering three-dimensional vascularized skeletal muscle constructs from myoblasts, fibroblasts and ECs. Furthermore, it was shown that prevascularization improves the in vivo performance of the tissue construct in three different models in mice (Jain et al., 2005). Koike et al.

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(2004) showed the formation of a network of long-lasting blood vessels in mice by co-implantation of vascular ECs and mesenchymal precursor cells in a collagen scaffold. These networks were revealed to be stable and functional for one year in vivo. Another innovative approach to co-culture in liver regeneration is the use of double-layered cell sheets. Harimoto et al. (2002) used a thermo-responsive culture dish grafted with poly (N-isopropylacrylamide) to make two contiguous cell sheets, one of human aortic ECs and the other of hepatocytes. The two layers were combined, creating a double-layered co-culture that exhibited expression of differentiated functions of hepatocytes. Alternatively, regeneration of tissue can be achieved without scaffolding material. Kelm et al. (2006) developed a scaffold-free vascularized artificial macrotissue. They used customshaped agarose moulds and induced the assembly of monodispersed primary cells by gravity-enforcement, forming minimal tissue units in this way. Macrotissues, assembled from HUVECs coated with human myofibroblasts, developed a vascular system that functionally connected to the chicken embryo’s vasculature after implantation. Unger et al. (2007) showed that it is possible to generate a prevascularized microcapillary-like network in vitro on biomaterial prior to implantation. They established a co-culture system consisting of HDMEC and primary osteoblasts or the human osteoblast-like cell line in several materials, namely silk fibroin. They demonstrated the in vitro formation of microcapillary-like structures containing lumen. In the same work, Unger et al. (2007) also reported a very interesting fact, the extensive vascular-like network formed in co-culture did not require the exogenous supply of angiogenic factors, thus equating therapies involving the incorporation of pro-angiogenic delivery systems into biomaterial scaffolds for bone regeneration. Fuchs et al. (2007) investigated OECs in direct two-dimensional and three-dimensional co-culture systems with the cell line MG-63 and human primary osteoblasts. They reported the formation of highly organized microvessel-like structures, in contrast to HUVEC where these structures were not observed. In a more fundamental work, Wenger et al. (2004) introduced a culture model system to assess the mechanisms involved in regulating angiogenesis during bone formation. HUVECs were grown as three-dimensional multicellular spheroids and seeded in a collagen matrix where ECs formed tubular outgrowths. When human osteoblasts were incorporated into the EC spheroids, thus forming heterogeneous cospheroids, the ability of EC spheroids to form tube-like structures under angiogenic stimulation was suppressed. The authors explain this contradictory result with the fact that direct contact or close proximity between ECs and osteoblasts overrides any angiogenic stimulation provided by soluble angiogenic factors. Other researchers have also highlighted the importance of direct cell-cell contact between these two cell types. For instance, osteoblastic cell differentiation analysis performed

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using different co-culture models with direct contact revealed that alkaline phosphatase activity was only increased by direct contact with human osteoprogenitor cells with human primary vascular endothelial types. Connexin43, a specific gap junction protein, has been proposed as being involved in this heterotypic communication (Guillotin et al., 2004).

30.5.5 Microsurgery strategies A flap may be defined as a segment of tissue with an independent blood supply. These may be classified into local, regional and free flaps. Free flaps are areas of tissue with an inherent vascular network supplied by a single vascular pedicle (Kannan et al., 2005). The work of Warnke et al. (2004) in clinical studies gave some visibility to the use of free flaps for the vascularization of bone grafts. A mandibular defect was scanned with threedimensional computed tomography (CT), and a titanium mesh was created. This cage was filled with bone mineral blocks and the patient’s bone marrow, loaded with BMP7 and implanted in latissimus dorsi muscle for seven weeks, then finally transplanted as a free bone-muscle flap to repair the mandibular defect. There was new bone formation and the patient had a clear improvement in his quality of life (Warnke et al., 2006). Another microsurgery technique is the insertion of an arteriovenous loop around the graft. Kneser et al. (2006) induced axial vascularization in a processed bovine cancellous bone matrix using a microsurgically constructed arteriovenous loop. Lokmic et al. (2007) reported another vascularization model where an arteriovenous loop was placed in a noncollapsible space protected by a polycarbonate chamber, the idea being that the arteriovenous loop presents a temporal window for cellular manipulation between the angiogenic growth phases and commencement of remodelling of the formed tissue. This time window, between seven and ten days, presents the ideal opportunity for seeding with stem cells or progenitor cells that will develop into specific tissue types while simultaneously being nourished by the developing microcirculation. Nevertheless, it must be remembered that reconstructive surgery with free flaps and arteriovenous loops increases morbidity in the patient (Kannan et al., 2005) and therefore alternatives for reconstructing the capillary bed are needed.

30.6

In vivo models to evaluate angiogenesis in tissue engineered products

30.6.1 The chick chorioallantoic membrane (CAM) The chick chorioallantoic membrane (CAM) assay is probably the most widely used in vivo assay for studying angiogenesis (Staton et al., 2004).

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CAM is a transient, densely vascularized organ, located underneath the shell membrane and eggshell. Because of this anatomical localization, it is easily accessible for experimental studies (Hagedorn, 2004). The CAM system has been extensively used for research in tumour biology, particularly for the study of tumour invasion and metastasis. In the biomaterials field, it has been used to study neovascularization of implants and to evaluate the inflammatory potency of biomaterials (Klueh et al., 2003; Zwadlo-Klarwasser et al., 2001). Nevertheless there are some limitations associated with this assay. The formation of a secondary vasoproliferative response, as a consequence of non-specific inflammatory reactions, impedes the quantification of the primary response. Another drawback is that when test material is placed on existing vessels, newly formed blood vessels grow within the CAM mesenchyme and therefore it is harder to distinguish real vascularization from a falsely increased vascular density (Ribatti et al., 2001). Nonetheless, there are several advantages in the use of this assay, including the fact that it is cheaper, easier to use and has fewer ethical concerns than other in vivo models (Zwadlo-Klarwasser et al., 2001). Despite these advantages, there is still limited data on biomaterials, mainly from natural origin, from CAM assays.

30.6.2 The dorsal skinfold chamber The dorsal skinfold chamber is an ectopic model performed on rats and mice. In this model, a piece of skin is removed from an anaesthetized animal and the biomaterial is placed on the exposed surface and covered by glass, which is then secured in place. Once the animals have recovered, these models allow for the continuous measurement of various parameters in living animals, including gene expression, angiogenesis, pH and blood flow (Staton et al., 2004). Intravital microscopy is the method used to visualize in a direct, continuous and non-invasive way the microvasculature at the level of individual microvessels. Contrast enhancement with fluorescently labelled dextrans or albumin enables the visualization of angiogenic sprouts and individual microvessel (Farhadi, 2004). The major advantage of the dorsal skinfold preparation is that the microcirculation can be analyzed through the observation window repetitively in unanaesthetized animals over a period of three to four weeks. On the other hand, there are some limitations in the use of this chamber for studying angiogenesis in tissue engineering constructs. The size of the constructs should not exceed 5 mm in diameter (width and length) to adequately fit within the 11-mm-sized chamber. Moreover, the height of the construct should be limited to 1 mm to ensure adequate closure of the chamber tissue by the glass cover (Laschke et al., 2006). Despite these limitations, the dorsal skinfold chamber model is still an ideal tool for the long-term in vivo study of blood

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vessel growth and remodelling in biomedical materials used in the field of tissue engineering.

30.7

Future prospects

It is our deeply held belief that innovative strategies to increase vascularization of tissue engineering constructs will be developed within the next few years. Nevertheless, in strategies that include ex vivo tissue generation, transplanted cells will always face the transition from the in vitro environment, saturated with oxygen, to the hypoxic in vivo environment. One way to minimize this change in oxygen tension would be to adapt the cell construct to hypoxic conditions prior to transplantation. One future trend will therefore be the dynamic in vitro culture of cell constructs under low oxygen tension, as a method to not only accelerate vascularization, but also to adapt the cells to the oxygen levels they will face when implanted. Major advances in scaffold architecture will be achieved using innovative processing methodologies such as microfabrication. Regarding co-culture for bone regeneration, the trend will be towards more complex systems involving the simultaneous culture of three or more cell types, such as ECs, osteoblast and pericytes or smooth muscle cells. These last two cell types are fundamental for the stability of newly formed blood vessels.

30.8

Sources of further information and advice

The 2004 review by Muschler et al. illustrates well the problems of mass transport in engineered tissues. Endothelial Cell Biology gathers a series of protocols relating to ECs, especially isolation of macro- and microvascular ECs, precursor cells, characterization, and in vitro and in vivo functional assays (Augustin, 2004). Embryonic Stem Cells: A Practical Approach provides further protocols for the isolation of EPCs (Notarianni, 2006).

30.9

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