Natural-Based Polymers for Biomedical Applications

i Natural-based polymers for biomedical applications © 2008, Woodhead Publishing Limited ii Related titles: Biomed...

Author: R.L. Reis | N.M. Neves | Joao F. Mano | Manuela E. Gomes | Alexandra P. Marques | Helena S. Azevedo

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i

Natural-based polymers for biomedical applications

© 2008, Woodhead Publishing Limited

ii

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


iii

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


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


v

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


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

vi

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


106 106 107 111 120

Contents

vii

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


193

viii

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


291 291 293 298

Contents

ix

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


357

x

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


396 396 397 400 402 404 405 409 409

Contents

16

Polyhydroxyalkanoate and its potential for biomedical applications

xi

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


485

487

xii

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


549

Contents

21.2

xiii

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


624 625

xiv

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


699

Contents

xv

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


761 761 762 763

xvi

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


765 774 776 776 776

xvii


(* = 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]

xviii


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



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


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]


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]



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


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


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]


xx


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]


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]


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]


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



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


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




xxii


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



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]


xxiii

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


xxiv

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


Preface

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


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

4


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 intermolecular 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


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


6


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


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


8

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


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.


10

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



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-


12


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.


OH

H

H

OH

COOH

COOCH3 O H H H

OH H H

OH

O H

O H H O OH H H

OH


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.


14


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



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,


16


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


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.


17

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


OH OH

O

OH

OH


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


20


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 ( 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



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


22


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



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).


24


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



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,


26


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



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


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.



29

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.


30


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,



31

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


32

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

(m250 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.


70


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



71

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


72


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,



73

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,


74


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



75

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.


76

2.8.2


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



77

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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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.


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).


90


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



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).


SPCL

91

92


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



93

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


94


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 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.



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.


Demineralization

Decolorization

132


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.


NH2

H

H

OH

O

H CH2OH

O

n


133

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


134


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



135

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


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.


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):


138


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



139

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


140


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


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.



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.



5.5

145

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



147

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.

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Part II Surface modification and biomimetic coatings


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]

•


Irradiation (gamma, UV, laser)

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


167

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


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:


172

H O

HO

H

O

H

HO HO

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


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


H .. O


173

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.


Temperature

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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


176


(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


178

(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


100 µm


179

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.


180


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


182

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).



OH


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-



185

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.


186


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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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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7.2


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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7.3.1


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


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.


Understanding molecular activation

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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).



7.4.2

211

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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Natural-based polymers for biomedical applications (e)

(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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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.


O


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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

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8

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8

(a)

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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)

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0 1

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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)


0.05M NaCl, pH = 4

Linear

0.1 M Acetate buffer pH 4

Linear

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8.2.2


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


242


(a)

(b)

(c)

(d)

(e)

(f)

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



243

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



245

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 (%)

80

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 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.,


Electrospun elastin and collagen nanofibers

321

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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(a)

(b)

(c)

(d)

(e)

(f)

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 (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


Polyhydroxyalkanoate and biomedical applications

429

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


430


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



431

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,


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


Contaminant

433


434

Natural-based polymers for biomedical applications 108 107

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



437

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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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

Electrospinning of natural proteins for tissue engineering

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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



449 (–)

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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 0.05). 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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17.5 Sizes of electrospun fibers at different delivery rates. (a) gelatin and collagen fibers; (b) elastin and tropoelastin fibers.

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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.,

(a)

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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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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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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

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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

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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

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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 +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.


Chitosan–polysaccharide blended nanoparticles

667

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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25.6


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

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Part VI Biocompatibility of natural-based polymers


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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26.2


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.



26.6

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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

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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.


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.



27.8

707

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.

27.10 References 1 Llull R (1999), Immune considerations in tissue engineering, Clin Plast Surg, 26, 549–568, vii–viii. 2 Tang L and Hu W (2005), Molecular determinants of biocompatibility, Expert Rev Med Devices, 2, 493–500. 3 Kirkpatrick C J, Krump-Konvalinkova V, Unger R E, Bittinger F, Otto M and Peters K (2002), Tissue response and biomaterial integration: the efficacy of in vitro methods, Biomol Eng, 19, 211–217. 4 Gretzer C, Emanuelsson L, Liljensten E and Thomsen P (2006), The inflammatory cell influx and cytokines changes during transition from acute inflammation



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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).



28.5

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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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28.6


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

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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


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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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.,


O2 and nutrients


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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


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

References

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