Tissue Engineering and Biodegradable Equivalents: Scientific and Clinical Applications

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I -issue bngineering

and Biodegradable Equivalents Scientific and Clinical Applications edited by

Kai-U we Lewand rowski Massachusetts General Hospital Boston, Massachusetts, U.S.A.

Donald L. Wise, Debra J. Trantolo, Joseph D. Gresser Cambridge Scientific, Inc. Cambridge, Massachusetts, U.S.A.

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Michael J Yaszemski The Mayo Clinic Rochester, Minnesota, U.S.A.

David E. Altobelli Harvard Dental School Nashua, New Hampshire, U.S.A.

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MARCELDEKKER, INC. ~

D E K K E R

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NEWYORK BASEL

ISBN: 0-8247-0755-9 This book is printed on acid-free paper. Headquarters Marcel Dekker, Inc. 270 Madison Avenue, New York, NY 10016 tel: 212-696-9000; fax: 212-685-4540 Eastern Hemisphere Distribution Marcel Dekker AG Hutgasse 4, Postfach 812, CH-4001 Basel, Switzerland tel: 41-61-261-8482; fax: 41-61-261-8896 World Wide Web http://www.dekker.com The publisher offers discounts on this book when ordered in bulk quantities. For more information, write to Special Sales/Professional Marketing at the headquarters address above. Copyright © 2002 by Marcel Dekker, Inc. All Rights Reserved. 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 and retrieval system, without permission in writing from the publisher. Current printing (last digit): 10 9 8 7 6 5 4 3 2 1 PRINTED IN THE UNITED STATES OF AMERICA

Preface

Tissue engineering involves the development of a new generation of materials or devices capable of specific interactions with biological tissues. These combine novel materials with living cells to yield functional tissue equivalents. Such systems are useful for organ tissue replacement where there is a limited availability of donor organs or where (e.g., for nerves) no natural replacements are available. These constructs are also useful as delivery vehicles for gene therapy. Controlling the tissue structure surrounding or within a material construct ultimately requires control of the cells of the host that are drawn to or are affected by the implant. Tissue engineering is built upon the basic cell biology of these host cells and the variety of signals that control their behavior. Materials are a key ingredient in tissue engineering applications. Development of these materials is in a constant state of activity; the challenge is to replace old materials with new ones that allow better exploitation of advances in a number of technologies, such as drug delivery, recombinant DNA techniques, bioreactors, stem cell isolation, cell encapsulation and immobilization, and 2D and 3D scaffolds for cells. These newer materials are coming to the forefront in modern applications in the emerging world of tissue engineering. This new reference text, Tissue Engineering and Biodegradable Equivalents: Scientific and Clinical Applications, focuses on materials used for the regeneration of tissues in the human body. Chapters deal with issues in the selection of proper biomaterials from biocompatibility to biostability to structure/function relationships. Chapters also focus on the use of specific biomaterials based on their physiochemical and mechanical characterizations. Integral to these chapters are discussions of standards in analytical methodology and quality control. The users of Tissue Engineering and Biodegradable Equivalents will have a broad base of backgrounds ranging from the basic sciences (e.g., polymer chemistry and biochemistry) to more applied disciplines (e.g., mechanical and chemical engineering, orthopedics, and pharmaceutics). To meet varied needs, each chapter provides clear and fully detailed discussions. This in-depth, but practical, coverage should also assist recent inductees iii

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to the biomaterials circle. We trust that this reference textbook conveys the intensity of this fast-moving field in an enthusiastic presentation. Kai-Uwe Lewandrowski Donald L. Wise Debra J. Trantolo Joseph D. Gresser Michael J. Yaszemski David E. Altobelli

Contents

Preface Part I: Material Considerations in Tissue Engineering 1. Biomaterials for Tissue Engineering Jennifer Hartt Elisseeff, Robert Langer, and Yoshihiko Yamada 2. Fundamental Physiological Factors Directing Bone Tissue Engineering Design and Development Kacey G. Marra, Phil G. Campbell, Yunhua Hu, and Jeffrey O. Hollinger 3. Mimicking the Natural Tissue Environment Christopher J. Woolverton, Judith A. Fulton, Stephanie T. Lopina, and William J. Landis 4. Biocompatibility, Biostability, and Functional Structural Relationships of Biomaterials E. G. Nordström 5. Biodegradable Hybrid Porous Biomaterials for Tissue Engineering Tetsuya Tateishi, Guoping Chen, Takashi Ushida, Toshimi Murata, and Shuichi Mizuno 6. Lactide Copolymers for Scaffolds in Tissue Engineering Shin-Ichiro Morita and Yoshito Ikada 7. Biodegradable Urethanes for Biomedical Applications Sudha Agarwal, Robert Gassner, Nicholas P. Piesco, and Sudhakar R. Ganta 8. Significance of Drug Delivery in Tissue Engineering Yoshito Ikada and Yasuhiko Tabata 9. Electrospinning of Polymer Scaffolds for Tissue Engineering Gary L. Bowlin, Kristin J. Pawlowski, Joel D. Stitzel, Eugene D. Boland, David G. Simpson, John B. Fenn, and Gary E. Wnek 10. Clinical and Biomechanical Design Considerations for Engineered Cortical Bone Allografts Kai-Uwe Lewandrowski, Francis J. Hornicek, Frank X. Pedlow, Mark C. Gebhardt, Henry J. Mankin, and William W. Tomford

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Part II: Tissue Engineered Cartilaginous Materials 11. Material Selection for Engineering Cartilage Giuseppe M. Peretti, Jian-Wei Xu, and Mark A. Randolph 12. Biodegradable Scaffolds for Meniscus Tissue Engineering Mark A. Sweigart and Kyriacos A. Athanasiou 13. Type I Collagen–Based Template for Meniscus Regeneration Shu-Tung Li, William G. Rodkey, Debbie Yuen, Peggy Hansen, and J. Richard Steadman 14. Biomaterials for Cartilage Tissue Engineering Hani A. Awad, Geoffrey R. Erickson, and Farshid Guilak Part III: Bone Repair Biomaterials 15. Polymeric Biodegradable Hard Tissue Engineering Applications Gamze Torun Köse and Vasif Hasirci 16. Controlled Porosity of Tissue-Engineered Cortical Bone Grafts Kai-Uwe Lewandrowski, Shrikar P. Bondre, Debra J. Trantolo, and Donald L. Wise 17. Biodegradable Scaffolds as Bone Graft Extender Kai-Uwe Lewandrowski, Shrikar P. Bondre, Debra J. Trantolo, and Donald L. Wise 18. Bioactivity of Nanohydroxyapatite in a Scaffold for Periodontal Repair Kai-Uwe Lewandrowski, Shrikar P. Bondre, Donald L. Wise, Jacqueline Y. Ying, and Debra J. Trantolo 19. Segmental Bone/Joint Replacement Using Guided Tissue Regeneration Edmund Y. Chao, Nozomu Inoue, Frank J. Frassica, and Frank H. Sim 20. Injectable Calcium Phosphate Cements for Repair of Bone Defects Shigeo Niwa and Racquel Z. LeGeros 21. Inorganic Bone Substitutes R. Schnettler and E. Dingeldein 22. Demineralization and Perforation of Cortical Bone Allografts: Preparatory Methods Kai-Uwe Lewandrowski, William W. Tomford, and Henry J. Mankin Part IV: Gene Therapy Applications 23. Adenovirus Vector–Mediated Gene Transduction for the Treatment of Bone and Joint Destruction of Rheumatoid Arthritis Sakae Tanaka 24. Muscle-Derived Cell–Based Gene Therapy and Tissue Engineering for the Musculoskeletal System Nobuo Adachi, Dalip Pelinkovic, Kenji Sato, Freddie H. Fu, and Johnny Huard Part V: Engineering of Muscle and Skin Biometerials 25. New Methods to Enhance the Regeneration of Muscle Jacques Ménétrey, Laurent Bernheim, Charles R. Bader, and Johnny Huard 26. Tissue Engineered Skin Ysabel M. Bello, Anna F. Falabella, and Robert S. Kirsner

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27. The Use of Muscle-Derived Cells for the Treatment of Muscle Pathologies Daniel Skuk and Jacques P. Tremblay 28. Soft Tissue Engineering with Ophthalmological Applications A. Tezcaner and Vasif Hasirci Part VI: Biomaterials in Urological Applications 29. Biomaterials for Tissue Engineering in Urology Byung-Soo Kim and Anthony Atala 30. Urologic Applications of Fibrin Sealant and Bandage R. Clayton McDonough III and Allen F. Morey Part VII: Sealants and Adhesives in Tissue Engineering 31. Evaluation of Biodegradable Fleece-Bound Sealing: History, Material Science, and Clinical Application Roman T. Carbon 32. Clinical Indications for Surgical Tissue Adhesives William D. Spotnitz, Sandra Burks, and David Mercer 33. Fibrin Glue Use in Surgery Peter P. Lopez, Qammar N. Rashid, and Stephen M. Cohn 34. Prevention of Cardiac Adhesions with the Use of Tissue-Engineered Biomaterials Arthur C. Hill and Trudy D. Estridge Part VIII: Analytical Tools for Biomaterials in Tissue Engineering 35. Combinatorial Cell Culture Applications to Tissue Engineering Joel S. Greenberger, Julie Goff, Michael W. Epperly, Alfred Bahnson, Douglas Koebler, Donna S. Shields, Johnny Huard, Karen Yanez-Hanley, and Raymond K. Hovck 36. Bioactive Extracellular Matrices: Biological and Biochemical Evaluation Andrea Liebmann-Vinson, John J. Hemperly, Richard D. Guarino, C. A. Spargo, and M. A. Heidaran 37. Biomechanical Comparison of Biodegradable Lumbar Interbody Fusion Cages F. Kandziora, R. Pflugmacher, R. Kleemann, Kai-Uwe Lewandrowski, Donald L. Wise, and G. Duda 38. Structural, Chemical, and Mechanical Characterization of the Dentin/Adhesive Interface J. Lawrence Katz, Paulette Spencer, Yong Wang, Ajay Wagh, Tsutomu Nomura, and Sauwanan Bumrerraj Index

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1 Biomaterials for Tissue Engineering Jennifer Hartt Elisseeff* and Yoshihiko Yamada National Institute of Dental and Craniofacial Research, Bethesda, Maryland Robert Langer Massachusetts Institute of Technology, Cambridge, Massachusetts

I INTRODUCTION Tissue engineering is an interdisciplinary field that incorporates principles of engineering with the life sciences for the development of tissue or organ replacements. The fields of polymer chemistry; materials science; chemical engineering; and cellular, molecular, and developmental biology may all be applied to tissue engineering, demonstrating the multidisciplinary approach that must be taken to solve the problem of tissue and organ replacement. Three general strategies for tissue engineering have been adopted and focus on [1] 1. Isolated cells. This technique allows tissue replacement using cells that supply a specific function. 2. Tissue-inducing substances. This approach includes using growth factor delivery, gene therapy, or other methods to induce tissue formation or organ development. 3. Cells placed on or encapsulated within matrices. This method for tissue engineering incorporates synthetic or natural materials as a scaffold for tissue development. This technique allows for three-dimensional tissue development, maintenance of tissue shape, and incorporation of signal molecules and gene delivery. This chapter surveys synthetic and natural polymers used in tissue engineering and discusses methods for processing, characterizing, and testing biomaterials. Potential clinical applications are presented throughout.

* Current affiliation: Johns Hopkins University, Baltimore, Maryland

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II SYNTHETIC POLYMERS Synthetic biomaterials are continually being developed for tissue engineering and other purposes. Manipulation of polymer chemistry provides control of features such as the rate and mechanism of degradation, hydrophilicity/phobicity, swelling, and mechanical strength. These features may be regulated in synthetic biomaterials by polymer backbone and side group chemistry, polymer molecular weight, or hydrogel pore size. Both synthetic and natural biomaterials may be solid (not water-soluble) scaffolds or water-soluble polymers capable of forming polymer networks that absorb water. Non–water-soluble polymers are processed using various methods to produce a wide range of scaffolds, such as felts, sponges, tubes, or fibers, that attempt to organize cells three-dimensionally and control cell function (see Section IV). Hydrogels offer advantages of in situ formation and cell encapsulation, as presented in Section III. For many tissue engineering applications, one polymer cannot provide all of the desired biological and physical properties. Thus, copolymers containing multiple monomers with different chemistries are synthesized or a scaffold may contain more than one (co)polymer to create a polymer blend. Copolymers and polymer blends allow endless possibilities for biomaterial and scaffold design. Table 1 outlines selected synthetic polymers and their potential clinical applications. A Linear Aliphatic Polyesters Linear aliphatic polyesters are a group of synthetic biodegradable polymers that are extensively used in medicine and in the field of tissue engineering [2,3]. Linear aliphatic polyesters include polyglycolide (PGA), polylactide (PLA), polycaprolactone (PCL), polyhydroxybutyrate, and their copolymers (e.g., polylactide-co-glycolide). Figure 1 provides the chemical structures of selected polyesters. Polyesters, such as PGA, may be synthesized by direct condensation [for polymers with a molecular weight (MW) less than 10,000] or more commonly by a ring-opening polymerization of the cyclic dimer [3,4]. Both PGA and PLA have been used extensively in medicine for applications such as wound closures (e.g., Dexon®, Vicryl®). These polymers have also had an important impact in the field of tissue engineering, demonstrating their ability for supporting cell proliferation and differentiation for a variety of cell types and tissues.

Table 1 Examples of Synthetic Polymers with Tissue Engineering Applications Polymer

Physical characteristics

Poly(esters) Poly(glycolic acid) Poly(lactic acid) Poly(caprolactone)

Solid: fiber, tube, sponge, screw, etc.

Poly(anhydride)

Solid, crosslinked network Solid, copolymer hydrogel Solid sponge, hydrogel Hydrogel, coating, solid

Poly(propylene fumarate) Poly(vinyl alcohol) Poly(ethylene glycol)

Potential clinical applications

Refs.

Cartilage, bone, muscle, nerve, blood vessel, valves, bladder, drug delivery, liver, cardiac tissue Bone, drug delivery

13, 14, 17, 18, 20, 21, 26, 31, 33, 36, 69 6, 47, 48

Bone, cardiovascular

60, 78

Cartilage, nerve Cartilage

76, 77 55, 73

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Figure 1 Chemical structure of poly(esters) commonly used for tissue engineering. The degradation properties of PGA and PLA have been extensively studied [5]. Degradation occurs by passive hydrolysis and a bulk degradative mechanism. Bulk degradation is characterized by a loss in polymer molecular weight while mass is maintained. Mass maintenance is useful for tissue engineering applications that require specific shapes, such as auricular or craniofacial augmentations in plastic surgery. Unfortunately loss in molecular weight causes a significant decrease in mechanical properties. Thus, polyesters may have significant mechanical strength initially, but bulk degradation may cause sudden loss in strength that can result in failure when used in orthopedic applications [6]. Degradation is dependent on water uptake, pH, crystallinity, steric hindrance, molecular weight, thermal history, and porosity [3]. Copolymers may be synthesized that allow degradation control on the order of weeks to years. The degradation products of PLA and PGA are lactic acid and glycolic acid. These products decrease the pH in the surrounding tissue, resulting in inflammation and potentially poor tissue development [7]. The polyesters PGA and PLA have end groups available for functionalization (i.e., addition of moieties to control biological and/or physical properties of biomaterials). Additional functional sites have been added to polyesters by copolymerization with molecules that add side chains and functional groups such as amino acids [8]. Poly(lactic acid-co-lysine-co-aspartic acid) (PLAL-ASP) was synthesized and functionalized with methacrylate

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groups to create a degradable PLA polymer network with aspartic acid side chains available for further functionalization [9]. PLA-PEO copolymers were synthesized to have the degradative and mechanical properties of PLA and the biological control provided with PEO and its functionalization [10,11]. Furthermore, PLA hydrophobicity and PEO hydrophilicity provide unique phase separation and physical characteristics [12]. B Clinical Applications Tissue engineering applications of polyesters are wide ranging and varied. PGA and PLGA copolymers have been applied to tissue regeneration in the vascular system. Niklason designed PGA tubes seeded with endothelial and smooth muscle cells for tissue engineering small diameter blood vessels [13]. The resulting blood vessels withstood physiological pressures and responded to vasoactive stimuli in vitro and remained patent in vivo in swine. Shum-Tim engineered aorta, a large diameter vessel, using a new polyester copolymer based on PGA and polyhydroxyalkanoate [14]. The aortas remained patent for up to 5 months compared to control constructs that were not seeded with cells and became occluded. Ratcliffe has reviewed these and other techniques for vascular tissue engineering including completely cellular systems, polyurethane, and collagen gels [15,16]. Other members of the vascular system that have been engineered include heart valves using polyhydroxyalkanoate, microvessels engineered on collagen gels, and cardiac tissue seeded on PGA felts [17–20]. Polyesters and other polymers have been applied to neural tissue engineering in place of silicon tubes and native vessels, previously used for nerve regeneration. Hadlock engineered PLGA foams with varying numbers of channels and observed axonal regeneration [21]. Schmidt demonstrated enhanced neurite outgrowth with electrical stimulation by the conducting polymer polypyrrole [22]. In addition, N-(2-hydroxypropyl) methacrylamide and collagen–glycosaminoglycan matrices have been seeded with neuronal and glial cells and found to stimulate axonal regeneration [23–25]. Many applications of tissue engineering have been studied in the musculoskeletal system including cartilage, bone, and muscle [26–30]. Again PGA meshes have been utilized for cartilage tissue engineering both in vitro in bioreactors and in vivo in rabbits [7,31]. Meniscus-like tissue has also been regenerated using PGA meshes and type I and II collagen gels [32,33]. Inorganic scaffolds such as hydroxyapatite and hydroxyapatite composites have played a significant role in bone tissue engineering [30,34,35]. Further discussion of musculoskeletal tissue engineering is provided in subsequent chapters. Polyesters have proven useful in urogenital tissue engineering applications. Overpenning used bladder-shaped PGA scaffolds seeded with urothelial and muscle cells to regenerate bladders in dogs [36]. Tissue engineered bladders demonstrated a urine capacity 95% of precystectomy volume compared to 46% with polymer alone and 20% in dogs that did not receive a graft. Additional applications of tissue engineering in urology are reviewed by Desgrandchamps [37] and in Part VI of this text. Polyglycolide has also been used to form soft tissue replacements and engineer oral tissues. Dental pulp has been engineered using PGA fibers, collagen I, and alginate gels [38,39]. Oral mucosa was synthesized using cadaveric dermis [40]. The resulting tissue had a well-stratified prekeratinized mucosa. While the tissue had many histological and biochemical similarities to native tissue, cells were in a hyperproliferative state and the fatty acid composition differed from native oral mucosa. A PLGA copolymer was used as a stent for tracheal reconstruction [41]. No cells were seeded on the stent, and no significant pre-

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vention in airway obstruction was observed. Small intestine and corneas have also been regenerated [42,43]. Regenerated small intestine stained positive for enzyme production similar to native tissue. Human corneal epithelial cells and fibroblasts seeded on collagen gels formed a basement membrane after only 3 days of incubation. The range of potential clinical applications for tissue engineering is wide, as demonstrated by the range of cells seeded or encapsulated in biomaterials. Tissue engineering has the potential to create a profound impact in clinical practice for organ and tissue replacement. C Polyanhydrides Polyanhydrides are another class of polymers used in tissue engineering and drug delivery. These polymers are synthesized by reaction of diacids with anhydride to form acetyl anhydride prepolymer. High molecular weight anhydrides are synthesized from the anhydride prepolymer in a melt condensation. Imides were copolymerized with polyanhydrides to create a polymer with increased mechanical properties while maintaining degradability via the anhydride units. Incorporation of the aromatic imide monomers to increase mechanical properties led to the use of polyanhydride-co-imides in orthopedic applications such as bone tissue engineering [6]. Anseth and coworkers substituted methacrylate groups on polyanhydrides to create photocrosslinked networks with controlled and predictable degradation and enhanced mechanical properties for orthopedic tissue engineering applications [44–46]. The general chemical structures of poly(anhydrides) and poly(imides) are given in Fig. 2. Polyanhydrides degrade by surface erosion in a highly predictable and controllable manner, making these polymers useful for drug delivery [47,48]. Surface erosion provides a more constant change in mass loss and mechanical property decrease compared to degradation that occurs by a bulk mechanism. Degradation in anhydride-based copolymers can be controlled by varying polymer backbone chemistry and monomer ratio.

Figure 2 General chemical structure of poly(anhydrides) and poly(imides).

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Figure 3 General chemical structure of poly(phosphazenes). D Polyphosphazenes Polyphosphazenes are polymers with a backbone containing a nitrogen and phosphorus atom, as depicted in Fig. 3. Polyphosphazenes have side groups that allow the synthesis of block and star polymers. In addition to the potential incorporation of biological moieties, side groups give control of polymer physical properties. For example, degradation can be altered by controlling the proportion of hydrolytically labile side groups. Side group chemistry allows synthesis of hydrophobic elastomers or hydrophilic hydrogels. Cytocompatibility of polyphosphazenes was studied for applications in skeletal tissue engineering [49,50]. Osteoblasts seeded on polyphosphazene sponges proliferated over the 1 week studied. III HYDROGELS Hydrogels are biomaterials that provide a number of advantages for tissue engineering and cell immobilization [51,52]. Hydrogels provide the possibility of minimally invasive implantation or injection of scaffold/cell constructs, high tissue-like water contents, and elasticity [53]. The ability to encapsulate or immobilize cells has made hydrogels particularly useful for tissue engineering. Figure 4 demonstrates cells encapsulated in a poly(ethylene glycol) and alginate hydrogel, respectively. Hydrogels are formed by crosslinking water-soluble polymers to form insoluble polymer networks. The networks are often capable of absorbing large amounts of water. Figure 5 demonstrates how individual polymer chains are crosslinked to form a network. Crosslinks may be formed by physical (hydrogen bonds, van der Waal interactions), ionic, or covalent bonds [54]. Formation of crosslinks may be triggered by temperature change (thermosets), ionic changes, addition of chemical crosslinkers (e.g., carbodi-mides, gluteraldehyde), or radiation exposure [52]. We have examined radiation (light) exposure as a method to encapsulate chondrocytes in a hydrogel [55]. Photopolymerization provides temporal and spatial control over gel formation both in vitro and in vivo. Enhanced control over polymerization using light facilitates in situ gelation and the development of minimally invasive techniques for scaffold implantation such as transdermal photopolymerization [73]. Photopolymerization requires addition of a photoinitiator to absorb light and to start a radical polymerization, a concern for cell biocompatibility and encapsulation. Cytocompatibility of various photoinitiators has been studied to develop photopolymerizing systems that are biocompatible for cell encapsulation [56]. Hydrogels may vary widely in their physical or structural characteristics, charges, pore size, and mechanical properties. One important property of hydrogels that can be

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Figure 4 Chondrocytes encapsulated in alginate (left) and poly(ethylene oxide) (right) show uniform dispersion in the hydrogel matrix.

manipulated is the pore size. The pore size affects the water content of the network, which is important for cell nutrient and waste transport and for the mechanical properties of the hydrogel [57]. The pore size, or average molecular weight between crosslinks, is characterized by measuring the swelling or mechanical properties of the gel. The equilibrium swelling volume, Q, is determined by the following equation: (p  s) Q   p where s and p are the volume of the solvent and polymer, respectively. A weight degree of swelling is an alternative method to determine the solute (sol) fraction of a hydrogel [58]. In either case the weight of the swollen and dried hydrogel must be obtained experimentally. From the equilibrium swelling volume, the average molecular weight between crosslinks can be calculated using the Peppas–Merrill equation. The compressive modulus, calculated from stress–strain curves, is correlated to the average molecular weight between crosslinks through the rubber elastic theory [52,59,60]. A Poly(Ethylene Glycol) Poly(ethylene glycol) (PEG) has an extensive history in biomedical applications. Because it offers limited protein and cell adhesion, immunogenicity, and antigenicity, PEG is often

Figure 5 Polymer networks or hydrogels are formed by crosslinking polymer chains. Crosslinks are created through physical interactions, ionic interactions, or covalent bonds.

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(a)

(b) Figure 6 Chemical structure of (a) poly(ethylene glycol) and (b) poly(ethylene glycol) functionalized with degradable polyester units.

used to coat medical devices to prevent protein and cell adhesion [61]. Both PEG and poly(vinyl alcohol) networks can act as tissue or mucoadhesives, providing unique applications such as the prevention of postoperative seromas or tissue adhesions [62–64]. The chemical structure of PEG is shown in Fig. 6a. Glycol end groups are available for chemical modification. The unique properties of PEG described led to its use in both cellular and tissue engineering. Effects of adhesion peptides on cell activity have been studied using PEG [65], which is ideal for this application due to its biocompatibility and general inertness to cell and protein adhesion. Attachment of bioactive peptides to PEG allows analysis of cell response and optimization of peptide spacing and clustering for cell presentation [66,67]. Only specific responses of cells toward ligands and biological molecules synthetically attached to the polymer are detected. Drumheller and Hubbell used PEG modified with the Arg-Gly-Asp adhesion peptide to determine that the required spacing of the peptide to cause fibroblast spreading was 440 nm [68]. Griffith has also synthesized well-defined cell adhesive substrates for studying hepatocyte biology and tissue engineering using PEG [69]. Poly(ethylene glycol) is not degradable but has been modified to form degradable PEG hydrogels. For example, Sawheny attached PLA and PGA degradable units and endcapped with polymerizing (meth)acrylate groups to form hydrolytically degrading hydrogels (Fig. 6b) [70]. West formed PEG hydrogels using degradable crosslinks sensitive to enzymatic cleavage, allowing network degradation to be biologically controlled [71]. Poly(ethylene glycol) has demonstrated utility in cartilage tissue engineering. High molecular weight PEG has been used to engineer cartilage-like tissue in vivo [72]. Primary chondrocytes were mixed with 100,000 MW PEG and injected subcutaneously in nude mice. Neocartilage was produced after 2 weeks implantation. Furthermore, primary chondrocytes were encapsulated in PEG-based photopolymerizing hydrogels in vitro and in vivo to form neocartilage [55,73]. The biochemical content and mechanical properties of constructs increased over time. Proteoglycan content was similar to native cartilage while collagen content was significantly lower. A copolymer of PEG and poly(propylene oxide) (PPO) is the basis for Pluronics. Pluronics forms a thermosensitive gel by aggregation of

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hydrophobic regions of the copolymer PPO. Chondrocytes have been encapsulated in pluronics for tissue engineering of auricular structures [74]. Complex helical shapes were injected subcutaneously in swine, and the developing neocartilage maintained the original shape of the pluronics. B Poly(Vinyl Alcohol) The lack of available functional groups on PEG hydrogels led researchers to examine poly(vinyl alcohol) (PVA) as a potential hydrogel scaffold [75]. Poly(vinyl alcohol) has pendant alcohol groups that are available for attachment of biological molecules, and PVA hydrogels may be synthesized using crosslinkers (e.g., gluteraldehyde), formation of physical crosslinks in crystalline regions, or by introduction of acrylate groups [63]. For example, PVA modified with varying amounts of acrylate groups was synthesized, and the resulting swelling and mechanical properties were studied [76]. Degradation rate and mechanical strength were modulated by the number of crosslinks in the hydrogel. Chondrocytes survived encapsulation in the PVA gels and produced proteoglycans and collagen, products of differentiated chondrocytes [76]. Poly(vinyl alcohol) may also be processed into foams on which cells may be seeded. Li seeded PC12 (pheochromacytoma) cells on PVA foams and observed efficient catecholamine release and cell distribution [77]. C Poly(Propylene Fumarate) Mikos and coworkers synthesized poly(propylene fumarate) (PPF) and poly(propylene fumarate) copolymers as a potential biomaterial [60]. Poly(propylene fumarate) is an unsaturated linear polyester (see Fig. 7 for chemical structure). It can be crosslinked with a vinyl monomer to create a rigid polymer network with significant mechanical strength for orthopedic applications. Degradation occurs by the hydrolytic mechanism described earlier for polyesters. Mechanical strength was also improved by the formation of PPF–-tricalcium phosphate composites [78]. Elasticity of PPF was increased for application as a cardiovascular stent by synthesis of the copolymer poly(propylene fumarate-co-ethylene glycol). The synthesis of PPF and PPF copolymers demonstrates how mechanical properties can be manipulated by polymer chemistry for diverse clinical applications. The polymers presented are examples of synthetic materials widely used for tissue engineering, but the list is by no means exhaustive. New materials are continually being synthesized. The synthesis of copolymers or polymer composites or blends may be necessary to engineer cell scaffolds with desired physical and biological properties for specific tissue engineering applications.

Figure 7 Chemical structure of poly(propylene fumarate).

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IV NATURAL SCAFFOLDS Many naturally occurring scaffolds have been investigated as biomaterials for tissue engineering purposes. The extracellular matrix is a very complex biomaterial that controls cell proliferation, migration, and differentiation leading to organ development and maintenance. In many ways, synthetic polymer scaffolds are designed to mimic specific functions of the extracellular matrix. Similar to the desired functions of synthetic tissue engineering scaffolds, the extracellular matrix interacts with cells and growth factors to control cell function and also provide mechanical integrity to the tissue [79]. Specific polymers found in the extracellular matrix have been examined for use in tissue engineering [80]. Natural polymers used for tissue engineering include proteins, poly(amino acids), and polysaccharides [81]. As in the case of synthetic scaffolds, natural scaffolds may be processed into foams or sponges or used as hydrogels to encapsulate cells. Natural or biological molecules, such as amino acids, are incorporated in synthetic scaffolds to create copolymers or composite scaffolds with natural and synthetic components. Collagen and fibrin glue will be presented as natural proteins, while alginate and chitosin will be discussed as polysaccharides for tissue engineering. Natural scaffolds and their potential applications are outlined in Table 2. A Collagen There are at least 22 types of collagen present in our bodies. Collagens I, II, and III may be considered classical collagens and are highly abundant in vivo [82]. The collagen chains form triple helices that are packed or processed into microfibrils. Collagen I is present in skin, tendon, cornea, dentin, and fascia, while collagen II makes up 10% wet weight of articular cartilage. Collagen II, a major collagen in cartilage, is assembled into fibrils with minor amounts of collagen IX and XI and crosslinked to form a scaffold that interacts with proteoglycan molecules and provides the impressive tensile strength that characterizes cartilage. Collagens can also promote cell adhesion as demonstrated by the Asp-Gly-Glu-Ala (DGEA) peptide in collagen I that functions as a cell binding domain [82]. While collagen has useful biological properties applicable to tissue engineering, collagen also offers flexibility in processing. Collagen can be processed into foams or sponges on which cells are seeded or gels capable of encapsulating cells. Collagen foams offer the advantage of pore structure control and a high surface area for cell adhesion. Spector synthesized foams of both collagen I and II, and copolymers with glycosaminoglycans (GAG) using a freeze-drying technique and additional ultraviolet crosslinking to increase mechanical integrity and slow degradation [83–85]. Canine chondrocytes have been seeded on col-

Table 2 Examples of Natural Polymers Used in Tissue Engineering Polymer

Physical characteristics

Potential clinical applications

Collagen

Solid sponges, gels

Fibrin Alginate

Hydrogel Hydrogel

Cartilage, tendon, skin, vessels, salivary gland, nerve, tendon, meniscus, cornea, kidney Cartilage, sealent Cartilage, muscle, dental pulp, soft tissue

Refs. 19, 23, 24, 32, 43, 57, 83–86, 88 102–105, 113 38, 74, 90–97

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lagen foams and incubated in vitro and implanted in vivo in dogs. Hyaline, articular, and fibrocartilage were present in the defects after 15 weeks. The regenerated tissue contained collagen II and proteoglycan, but did not attach well to the subchondral plate. Butler and Chamberlain engineered skin and nerve using collagen and other extracellular matrix-derived scaffolds [24,86]. Collagen can also be processed into a gel. Collagen gels are used for a wide variety of applications including studies on cell migration, differentiation, wound contraction, and renal tissue engineering [87,88]. Wound adhesive, or surgical glue, has been developed using collagen and poly(L-glutamic acid) with carbodiimide as a crosslinking agent [89]. The adhesive is applied to wounds in order to promote host regeneration of tissue on the collagen gel scaffold without scar formation. B Alginate Alginate is a polysaccharide found in seaweed that is composed of mannuronic and guluronic acid. The polysaccharide contains block copolymer regions rich in mannuronic or guluronic acid and random copolymer regions with random organization of the two sugars. Gelation occurs by interaction of divalent cations (e.g., Ca2) with blocks of guluronic acid from different polysaccharide chains. Cells can be encapsulated under physiological conditions and recovered by degelation of alginate using a chelating agent such as ethylene di-aminetetra-acetic acid (EDTA). The structure of the hydrogel may be controlled by varying the conditions of gelation. Aydelotte formed alginate spheres with channels in which chondrocytes from growth plate aligned in columns similar to their in vivo growth plate organization [90]. Unfortunately, the conditions required to form the alginate columns resulted in significant cell death. Due to the ionic nature of crosslinks in alginate gels, degradation or mechanical instability may occur. Rowley incorporated covalent crosslinks in alginate to address hydrogel stability [91]. Similar to PEG hydrogels, alginate does not interact with cells or proteins. Thus, Rowley also incorporated the cell adhesion peptide Arg-Gly-Asp to alginate in order to encapsulate anchorage-dependent cells such as myocytes. Examples of alginate processing conditions, chemical functionalization, and incorporation of biological peptides demonstrate the flexibility of natural scaffolds and their potential applications. Alginate has a long history for encapsulation of chondrocytes, an anchorage-independent cell type. Chondrocytes dedifferentiate into fibroblast-like cells when cultured in vitro on cell adhesive surfaces. Chondrocytes encapsulated in alginate remain differentiated with a spherical cell morphology and secrete the cartilage-specific markers collagen II and aggrecan. Encapsulation of chondrocytes in alginate allows monitoring of differentiated chondrocyte response to biological agents such as interleukin-1B in vitro [92]. Cellular activity and extracellular matrix production, quality, and mechanical integrity can be tested. Buschmann encapsulated chondrocytes in alginate and agarose (a thermoresponsive polysaccharide) and observed an increase in mechanical properties of the gels over time as extracellular matrix was produced [93]. Growth plate chondrocytes, fetal chondrocytes, and mesenchymal stem cells have been encapsulated in alginate [92,94,95]. In each of these systems, the chondrocytes demonstrated a differentiated phenotype, producing extracellular matrix and retaining a cell morphology typical of chondrocytes. For specific clinical application, Paige used alginate matrix as an injectable tissue engineering scaffold for cartilage replacement [96]. In this alginate system, cartilage was formed in different shapes as it was injected and the resulting cartilage-like tissue retaining the same shape. Masuda has developed a method termed alginate recovered chondrocytes (ARC) for scaffold-free car-

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tilage tissue engineering using alginate. After initial cartilage matrix deposition, the alginate is removed, leaving only neocartilage ready for implantation [97]. Other polysaccharide polymers that have been applied to tissue engineering include chitosan, hyaluronic acid, and gelatin. These polysaccharides have been modified and studied as scaffolds for tissue engineering [98]. Madihally prepared chitosan-gelatin networks by polyelectrolyte complex formation, freeze-drying, and crosslinking with gluteraldehyde [99]. Pore size was modified in gelatin hydrogels and sponges by freezing and using ice crystals as porogens [100]. Hyaluronic acid has carboxylic acid moieties available for modification including addition of crosslinkers for hydrogel formation or biologically active molecules [101]. C Fibrin Fibrin is a biological polymer which functions as a hemostatic barrier to prevent bleeding and as a natural scaffold for fibroblasts to support wound healing. In vivo, fibrin polymerizes by the conversion of fibrinogen to fibrin monomer by thrombin. Clinically used fibrin sealant (or glue) is composed of fibrinogen, factor XIII, and thrombin in addition to antifibrinolytic agents. Fibrin sealant may be made from autologous blood samples or from recombinant proteins. Fibrin glue has been widely used in surgical and endoscopic procedures for such applications as sealing lung tears, cerebral spinal fluid (CSF) leaks, and bleeding ulcers in addition to prevention of wound dehiscence with sutures [102]. Fibrin glue has been applied to cartilage tissue engineering [103]. Auricular structures have been formed from fibrin/chondrocyte constructs. Biochemical and mechanical analysis demonstrated tissue with cartilage-like properties [104]. Schense seeded human myofibroblasts on fibrin gels to create a scaffold with uniform cell distribution and minimal toxicity for cardiovascular tissue engineering [105]. Schense also incorporated bioactive peptides in fibrin gels to stimulate neurite outgrowth in vitro [105]. Fibrin glue may find further use in tissue engineering, particularly in material composites where mechanical and degradation properties may be improved. Fibrin is further discussed in Sections V and VI. V PROCESSING Solid polymer matrices may be processed in a variety of techniques to produce scaffolds with specific architectures such as pores. Solvent casting, melt processing, and membrane lamination are examples of processing techniques [106]. Polymer foams are created with a wide variety of porosities using solvent casting and particulate leaching [106]. Solvent casting is performed by dissolving a polymer in a solvent, usually an organic solvent. A particulate, or porogen, that is not soluble in the organic solvent is suspended in the polymer solution. The resulting polymer/porogen solution is placed in a mold, and the organic solvent is evaporated. The particulate is subsequently dissolved using a solvent in which the polymer is not soluble. For example, PGA is dissolved in chloroform and a sugar is suspended in the solution. The chloroform is evaporated and the sugar is dissolved by washing with water to create a PGA scaffold with pores. Zhang melted sugars into fibers and thin discs to act as porogens and used the solvent casting technique to create PLA scaffolds with a nanofibrillar architecture [107]. Other porogens include salts and gelatin. Melt processing uses high temperatures to melt a polymer into a desired shape. Niklason formed tubular structures of PGA for blood vessel tissue engineering using melt processing [13].

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Finally, membrane lamination uses polymer dissolved in a solvent to glue sheets of polymer together [108]. Polymer processing allows great flexibility in scaffold design. Three-dimensional printing and other computer-assisted techniques have been designed to manufacture scaffolds with complex shapes on the order of nanometers. These complex structures have been designed to direct biological responses by mimicking natural structures such as capillary beds or extracellular matrix. Three-dimensional printing has been used with PGA and PLA to create complex structures for liver tissue engineering [69]. Cell adhesion has been selectively controlled on these complex structures by combining three-dimensional fabrication with surface modification using carbohydrate ligands [109]. Kaihara used silicon micromachining to engineer two-dimensional, micron-sized vascular channels [110]. Hepatocytes and endothelial cells seeded on the mold remained viable but were difficult to lift and transplant. Laser-based techniques have also been used with photopolymerizing polymers to create complex shapes from stored images [75]. Scaffolds for bone tissue engineering were created from magnetic resonance images (MRI) and CT images of craniofacial structures to manufacture patient-specific matrices [111]. Biological response to biomaterials may also be directed by surface chemistries and physical properties in addition to modification with biological peptides. Papadaki demonstrated the biological effect of varying polymer surface properties as measured by surface contact angle [112]. Skeletal muscle and chondrocytes demonstrated different rates of proliferation and differentiation as molecular weight and ratio of PEG and poly(butylene terephthalate) were varied in a polymer blend. Hydrogel physical structure can also be varied by altering the chemical composition and pore size. Dillon altered the average pore radii of agarose gels and synthesized agarose/chitosan and agarose/alginate composite hydrogels to study the influence of physical structure and charge on neurite extension [113]. The polycationic polysaccharide agarose/chitosan demonstrated the greatest neurite extension length in vitro when compared to the unmodified agarose and the negatively charged polyanionic agarose/alginate composite hydrogels which inhibited neurite extension. It is clear that polymer surface chemistry, physical properties, and biological modification must be studied for each specific tissue cell engineering application. VI POLYMER CHARACTERIZATION Biomaterials for tissue engineering must be characterized chemically, physically, and biologically. Nuclear magnetic resonance (NMR) and infrared spectroscopy (IR) are two common methods used to determine the chemical structure of a polymer. The structure of biological or natural polymers is often more complex than synthetic polymers and thus less characterized [79,114]. For example, Matrigel is a thermoset gel derived from the extracellular matrix of mouse tumors used to promote cell differentiation in vitro. While Matrigel is known to be rich in laminin and other basement membrane proteins, the exact composition of the gel is unknown. Synthetic polymers may be physically characterized with respect to their molecular weight, melting temperature (Tm), and glass transition temperature (Tg) [115]. Characterization of molecular weight and physical properties of a polymer are important for predicting polymer performance. The molecular weight of a polymer may be characterized using the number-average molecular weight (Mn), weight-average molecular weight (Mw), and polydispersity (PDI). Mn is determined experimentally by determining the number of polymer molecules in a sample of polymer and is defined as the total weight of all molecules in a polymer sample divided by the total number of moles present. Mw is the weight fraction

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of molecules of a specific molecular weight. The weight-average molecular weight is more biased toward higher molecular weights, thus MwMn. The PDI of a sample describes the range of molecular weights of polymer and is defined as Mw /Mn. Mw, Mn, and PDI can be practically estimated using gel permeation chromatography (GPC). Tm is the melting temperature of crystalline domains within a polymer sample. Tg is the temperature where amorphous domains of a polymer become brittle and rigid [116]. If a polymer sample has only crystalline domains, it will have a Tm and not a Tg, while a completely amorphous sample will have a Tg but no Tm. Tg and Tm are commonly determined experimentally using differential scanning calorimetry (DSC). As described in Section IV, polymers and polymer composites can be processed by a variety of methods. The resulting scaffold structure, pores, surface chemistry, and topology can be studied using techniques including electron microscopy, electron microspcopy for chemical analysis (ESCA), and scanning probe microscopies [117]. Mercury porosimetry is a classical method used for pore size and pore size distribution analysis. Atomic force microscopy (AFM) is one type of scanning probe microscopy for study of molecular and cellular level interactions at the biomaterial interface on a nanometer scale. For example, AFM has been used to monitor the surface properties of PLA/poly(sebacic acid) blends as degradation occurred. The AFM images demonstrated islands of PLA dispersed in the PSA with spherulites of PSA on the surface [118]. The distribution of polymers is an important component of biomaterial scaffold design since cellular interactions are affected. A review of scanning probe methodology is provided by Garrison and Ratner [119]. Cellular interactions with biomaterials may be evaluated using a variety of microscopy techniques, including optical, fluorescence, and newly developed techniques such as total internal fluorescence microscopy [10,120]. Mechanical characterization of biomaterials and engineered tissue constructs is critical to determining potential clinical use, particularly in orthopedic or vascular applications where mechanical integrity is critical. Understanding material stress and strain provides understanding of potential mechanical failure and fatigue. Changes in mechanical integrity with degradation should also be monitored [78]. The compressive, tensile, shear, and flexural moduli (static and dynamic) along with fracture strength are examples of mechanical characteristics that can be analyzed [121]. Most biomaterials may be characterized mechanically through compression tests, where a compressive modulus is determined from stress–strain curves, or by tensile tests. Biaxial mechanical testing was preferred to uniaxial compression or tensile testing for anisotropic biomaterials including small intestine submucosa. Small intestine submucosa is used as a repair biomaterial and as a tissue engineering scaffold [122]. Comprehensive hydrogel mechanical testing is reviewed by Anseth [59]. Developing tissues may require more specific tests to analyze mechanical function. For example, Bushmann monitored the evolution of mechanical properties of cartilagelike tissue in agarose gels using confined compression tests. Static and dynamic compressive strength and streaming potential were monitored as newly formed, mechanically functional extracellular matrix was synthesized in the agarose. The mechanical properties of the developing cartilage was directly related to the biochemical content of the developing matrix [93]. Furthermore, mechanical and biochemical properties of tissue engineered cartilage can be modulated by in vitro culture conditions to engineer cartilage with mechanical properties approaching native tissue [31]. Niklason measured the mechanical strength of developing vascular constructs and correlated the incremental modulus at high

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vessel wall stress with collagen content [13]. The incremental modulus was determined by varying static intraluminal pressure and measuring vessel response in digital images. Evaluating mechanical characteristics of a biomaterial and regenerated tissue is important for determining potential clinical application of materials and functionality of engineered tissues. Many biomaterials used for tissue engineering are degradable. Degradation is characterized both in vitro and in vivo. Polymer samples are generally incubated in vitro in buffer at 37°C. Buffer pH and sink conditions are maintained constant by periodic replacement. Polymer mass, molecular weight, shape, mechanical properties, and released degradation products can be monitored. Degradation is more difficult to monitor in vivo but the cage implant system discussed in Section VII allows some quantitation of degradation. Simple weight measurements and histological evaluation of shape may be performed to monitor degradation in vivo [78]. VII BIOCOMPATIBILITY The International Organization for Standardization (ISO) has developed guidelines for biological evaluation of medical devices that apply to biomaterials and tissue engineering systems [123]. First, a biomaterial or device must be tested for fitness for purpose. This includes studying the chemical, toxicological, physical, electrical, morphological, and mechanical properties. For example, the physical and biological requirements of a tissue engineering scaffold vary widely depending on the application. Orthopedic applications require scaffolds with high tensile and compressive strengths, while scaffolds used in the cardiovascular system require significant elasticity. The ISO guidelines for evaluation and testing for material selection include determining (1) materials required to manufacture the scaffold, (2) intended additives, process contaminants, and residues (e.g., initiator required for a polymerization), (3) leachable substances (important for in situ polymerizations where unreacted monomer is present), (4) properties and characteristics of the final product, and (5) degradation products. Evaluation of degradation products is particularly important for biodegradable materials in tissue engineering. Identification, quantification, and pharmacokinetic studies of the degradation process should be performed. A In Vitro Biocompatibility is tested in vitro and in vivo. Preferrably, biocompatibility should be tested in vitro as much as possible before being evaluated in vivo. For applications in tissue engineering, cytocompatibility of a biomaterial is particularly important since cells directly contact the material. In vitro compatibility is tested by exposing cells directly to a polymer surface, indirectly by exposing cells to polymer byproducts, or indirectly with an interposed layer between the cells and materials (i.e., agar diffusion test) [124]. Cell viability, proliferation, and functionality should be studied. Cell viability may be studied with vital stains such as the dyes neutral red and trypan blue. Living cells take up neutral red and exclude trypan blue such that live cells stain red and dead cells stain blue. MTT (3-[4-5dimethylthiazol-2-yl]-2-5-diphenyl bromide tetrazolium bromide) and alamar blue are two common methods to study cell toxicity and proliferation. MTT is reduced to blue formazan crystals by mitochondrial dehydrogenases and is measured spectrophotometrically [125,126]. MTT was used to visualize viable cells encapsulated in a poly(ethylene oxide)

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Figure 8 Viable chondrocytes encapsulated in a poly(ethylene glycol)–based hydrogel stain dark blue with MTT. hydrogel (Fig. 8). Alamar blue is a simple, one-step method to determine cell proliferation and toxicity that does not necessitate cell destruction [127]. Alamar blue is a nonfluorescent oxidized blue compound that is reduced to a pink fluorescent dye in viable cells. Finally, thymidine uptake is a method to measure cell proliferation by thymidine incorporation into DNA. B In Vivo In vivo compatibility is the next step for biomaterial compatibility testing [123]. According to the International Organization for Standardization, both long- and short-term effects of a biomaterial should be studied. Local reactions to the biomaterial and systemic effects should be studied. Effects include sensitization, irritation, hemocompatibility, genotoxicity, and carcinogenicity. Preliminary material testing in vivo is subcutaneous implantation, often in rats [128]. Acute and chronic inflammation is analyzed by implant analysis at 3-, 7-, 10-, and 14-day time points. Time points at 3 and 10 days measure peak acute and chronic inflammation, respectively. The implant, implant–cell interface, fibrous capsule, and surrounding tissue are analyzed. The implant chemistry, degradation, and mechanical strength are analyzed to observe the in vivo response of the biomaterial. Cell adhesion, particularly macrophages and other inflammatory cells, to biomaterials is analyzed on the implant interface. A fibrous capsule is usually formed around implants, with the exception of implants in bone. The thickness and biochemical composition of the capsule is another marker for measuring inflammatory response [128]. Finally, potential systemic effects must be analyzed through blood and urine sampling and determining the presence of any polymer byproducts in organs and tissues. Anderson developed the cage implant system to quantitatively and qualitatively monitor biomaterial inflammatory response [129]. The cage implant system offers the advantage of creating a chamber in which a material is implanted, and the inflammatory exudate can be periodically monitored without sacrificing the animal. Quantitative analysis of white cells, extracellular enzymes, and proteins in the exudate may be performed to quantitatively measure the extent of inflammatory response. Reviews of the cage implant system and in vivo

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compatibility further discuss the critical issue of determining in vivo inflammatory responses and compatibility of biomaterials [130]. C Cell Activity Cell differentiation is particularly important for tissue engineering applications since production of a specific extracellular matrix or cell activity is required for tissue development and maintenance. For example, collagen II and the proteoglycan aggrecan make up the majority of the extracellular matrix of cartilage. Thus, the production of collagen II and aggrecan are markers of chondrocyte differentiation that can be studied on a biomaterial. Cell proliferation may be similar on different polymer scaffolds, while expression of markers for cell differentiation vary. This was demonstrated by Calvert with bone marrow cells cultured on various poly(caprolactone) and PLGA surfaces [131]. Cell proliferation was similar on the different materials, but the expression of alkaline phosphatase and bone matrix mineralization was increased on the PCL and PLGA blends, suggesting their potential use over the homopolymers for bone tissue engineering. ACKNOWLEDGMENTS The authors would like to extend Hynda Kleinman and Joachim Seidel great appreciation for reviewing the manuscript. REFERENCES 1. Langer R., Vacanti J. 1993. Tissue engineering. Science 260:920–926. 2. Dunn, R. 1995. Clinical applications and update on the poly(alpha-hydroxy acids). In: Biomedical Applications of Synthetic Biodegradable Polymers. J. Hollinger, Ed. CRC Press: New York. 3. Behravesh E., Yasko A. W., Engel P. S., Mikos A. G. 1999. Synthetic biodegradable polymers for orthopaedic applications. Clin. Orthopaed. Rel. Res. 367S:S118–S125. 4. Peter S. J., Miller M. J., Yasko A. W., Yaszemski M. J., Mikos A. G. 1998. Polymer concepts in tissue engineering. J. Biomed. Mater. Res. 43(4):422–427. 5. Li S., Vert M. 1995. Degradable Polymers: Principles and Applications. Chapman and Hall: London. 6. Muggli D. S., Burkoth A. K., Anseth K. S. 1999. Crosslinked polyanhydrides for use in orthopedic applications: degradation behavior and mechanics. J. Biomed. Mater. Res. 46(2):271–278. 7. Freed L., Vunjak-Novakovic G., Biron R., Eagles D., Lesnoy D., Barlow S., Langer R. 1994. Biodegradable polymer scaffolds for tissue engineering. Bio/technology 12(7):689–693. 8. John G., Morita M. 1999. Synthesis and characterization of photo-cross-linked networks based on L-lactide/serine copolymers. Macromolecules 32(6):1853–1858. 9. Elisseeff J., Anseth K., Langer R., Hrkach J. 1996. Synthesis of photocrosslinked polymers based on poly(L-lactic acid-co-aspartic acid). Macromolecules 30:2182–2184. 10. Shakesheff K., Cannizzaro S., Langer R. 1998. Creating biomimetic micro-environments with synthetic polymer–peptide hybrid molecules. J. Biomater. Sci. Polym. Ed. 9(5):507–518. 11. Han D., Hubbell J. A. 1997. Synthesis of polymer network scaffolds from L-lactide and poly(ethylene glycol) and their interactions with cells. Macromolecules 30(20):6077–6083. 12. Kubies D., Rypacek F., Kovarova J., Lednicky F. 2000. Microdomain structure in polylactideblock-poly(ethylene oxide) copolymer films. Biomaterials 21(5):529–536. 13. Niklason L. E., Gao J., Abbott W. M., Hirschi K. K., Houser S., Marini R., Langer R. 1999. Functional arteries grown in vitro. Science 284(5413):489–493.

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Elisseeff et al. Transdermal photopolymerization of poly(ethylene oxide)–based injectable hydrogels for tissue-engineered cartilage. Plast. Reconstr. Surg. 104(4):1014–1022. Bryant S. J., Nuttelman C. R., Anseth K. S. 2000. Cytocompatibility of UV and visible light photoinitiating systems on cultured NIH/3T3 fibroblasts in vitro. J. Biomater. Sci. Polym. Ed. 11(5):439–457. Torres D. S., Freyman T. M., Yannas I. V., Spector M. 2000. Tendon cell contraction of collagen-GAG matrices in vitro: effect of cross-linking. Biomaterials 21(15):1607–1619. Brannon-Peppas L. 1994. Preparation and Characterization of Crosslinked Hydrophilic Networks. ACS: Washington, DC. Anseth K. S., Bowman C. N., Brannon-Peppas L. 1996. Mechanical properties of hydrogels and their experimental determination. Biomaterials 17(17):1647–1657. Suggs L. J., Kao E. Y., Palombo L. L., Krishnan R. S., Widmer M. S., Mikos A. G. 1998. Preparation and characterization of poly(propylene fumarate-co-ethylene glycol) hydrogels. J. Biomater. Sci. Polym. Ed. 9(7):653–666. Alcantar N. A., Aydil E. S., Israelachvili J. N. 2000. Polyethylene glycol–coated biocompatible surfaces. J. Biomed. Mater. Res. 51(3):343–351. Hill-West J., Chowdhury S., Sawhney A., Pathak C., Dunn R., Hubbell J. 1994. Prevention of postoperative adhesions in the rat by in situ photopolymerization of bioresorbable hydrogel barriers. Obstet. Gynecol. 83:59–64. Mongia N. K., Anseth K. S., Peppas N. A. 1996. Mucoadhesive poly(vinyl alcohol) hydrogels produced by freezing/thawing processes: applications in the development of wound healing systems. J. Biomater. Sci. Polym. Ed. 7(12):1055–1064. Sawhney A., Lyman F., Yao F., Levine M., Jarrett P. 1996. A Novel in situ formed hydrogel for use as a surgical sealent or barrier. In: 23rd International Symposium of Controlled Release of Bioactive Materials. Controlled Release Society: Kyoto, Japan. pp. 236–237. Drumheller P. D., Hubbell J. A. 1994. Polymer networks with grafted cell adhesion peptides for highly biospecific cell adhesive substrates. Anal. Biochem. 222(2):380–388. Kuhl P. R., Griffith-Cima L. 1996. Tethered epidermal growth factor as a paradigm for growth factor–induced simulation from the solid phase. Nat. Med. 2:1022. Maheshwari G., Brown G., Lauffenburger D. A., Wells A., Griffith L. G. 2000. Cell adhesion and motility depend on nanoscale RGD clustering. J. Cell Sci. 113(Pt 10):1677–1686. Schense J. C., Hubbell J. A. 2000. Three-dimensional migration of neurites is mediated by adhesion site density and affinity. J. Biol. Chem. 275(10):6813–6818. Kim S. S., Utsunomiya H., Koski J. A., Wu B. M., Cima M. J., Sohn J., Mukai K., Griffith L. G., Vacanti J. P. 1998. Survival and function of hepatocytes on a novel three-dimensional synthetic biodegradable polymer scaffold with an intrinsic network of channels. Ann. Surg. 228(1):8–13. Sawhney A., Pathak C., Hubbell J. 1993. Bioerodible hydrogels based on photopolymerized poly(ethylene glycol)-co-poly(-hydroxy acid) diacrylate macromers. Macromolecules 26:581–587. West J., Hubbell J. 1999. Polymeric biomaterial with degradation sites for proteases involved in cell migration. Macromolecules 32(1):241–244. Sims D., Butler P., Casanova R., Lee B., Randolph M., Lee W. P. A., Vacanti C., Yaremchuk M. 1996. Injectable cartilage using polyethylene oxide polymer substrates. Plast. Reconstr. Surg. 98:843–850. Elisseeff J., McIntosh W., Anseth K., Riley S., Ragan P., Langer R. 2000. Photoencapsulation of chondrocytes in poly(ethylene oxide)–based semi-interpenetrating networks. J. Biomed. Mater. Res. 51(2):164–171. Saim A. B., Cao Y., Weng Y., Chang C. N., Vacanti M. A., Vacanti C. A., Eavey R. D. 2000. Engineering autogenous cartilage in the shape of a helix using an injectable hydrogel scaffold. Laryngoscope 110(10 Pt 1): p. 1694–1697.

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2 Fundamental Physiological Factors Directing Bone Tissue Engineering Design and Development Kacey G. Marra, Phil G. Campbell, Yunhua Hu, and Jeffrey O. Hollinger Bone Tissue Engineering Center, Institute for Complex Engineered Systems, and Carnegie Mellon University, Pittsburgh, Pennsylvania

I INTRODUCTION Design and development of a bone tissue engineered structure must be directed by patient application. Bone tissue engineering deals with a structure that will be implanted into a patient. Therefore, the patient profile must provide the fundamental physiological guidance for the design and development pathway. Physiological patient considerations direct the pathway from laboratory to clinic. Consequently, merely addressing bone composition will result in a poorly designed product that will have an ineffective clinical outcome. The composition of bone is fundamental to bone tissue engineering. The “visibility” of composition has made it pre-eminent over other distinguishing parameters that must be considered. However, in terms of clinical practicality, this is a striking pitfall. Bone composition is invaluable for guiding design and development of tissue engineered bone when integrated with bone physiology. Composition provides the framework for the tissue engineer. Bone consists of organic and inorganic matrices. Several inclusive reviews referenced at the end of this chapter detail matrix composition and physical properties [1–6]. About 90% of the organic matrix is collagen, and the inorganic matrix is predominantly calcium and phosphate (i.e., calcium hydroxyapaptite). Dedicated to building and maintaining the matrices are three principle cell phenotypes: osteoblasts, osteocytes, and osteoclasts. At the end of this chapter references are provided for thorough reviews about bone cells [7–13]. The role of osteoblasts, osteocytes, and osteoclasts is to assist in calcium/phosphate balance and to sustain bone structure: the processes of remodeling. To fulfill these func25

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tions, a dedicated number of bone cells must be available. Cell populations are not static. Bone cells have a finite and limited life span, their renewal and replenishment is key to ensure balance in remodeling dynamics, specifically, an equivalent amount of bone resorbed is replaced. The product bone cells produce is sculpted into an integrated organic and inorganic matrix with a distinguished macroscopic appearance. Trabecular bone, spongy-looking in appearance, is a lacey network of interconnecting struts about 100–125 m in thickness and around 1–2 mm in length. Sandwiching trabecular bone is cortical bone. It appears denser and is formed by columns of Haversian systems, each about 200–300 m in diameter. Trabecular bone provides ready access to calcium and phosphate depots, thereby ensuring homeostasis. Osteoclastic/osteoblastic–molecular signaling controls cationic acquisition and disposition for homeostasis. In contrast, the less dynamic cortical bone provides biofunctional and locomotion roles, as well as protection for organs. Accomplishing homeostatic and biofunctional/locomotional roles requires trabecular and cortical bone to remodel: resorption and replacement. Approximately 20% of trabecular bone and 5% of cortical are remodeled each year. Osteoclasts resorb bone and osteoblasts replace it. Remodeling is a balanced process between these cells until the signals (e.g., hormonal) change and an imbalance occurs. With menopause there is a change and the signal (i.e., estrogen) imbalance profoundly affects osteoblasts. The consequence is resorption outpaces formation, bone mass plummets, and the disease osteoporosis results. Osteoblasts build bone; organic and inorganic matrices as well recruit and provoke lineage progression for osteoblast renewal. About 30–50% of osteoblasts undergo apoptosis (i.e., programmed cell death) every 3–4 weeks and must be renewed. Remaining osteoblasts become either osteocytes or lining cells. Osteoblast replenishment is accomplished by stem cell progression through a lineage pathway leading to the osteoblast. Stem cells are undifferentiated, and cues for their progression are molecular signals. If the signals (e.g., bone morphogenetic proteins) are sufficient and if the stem cells are responsive, osteoblasts develop. (Later in this chapter, the process will be described.) Aging and osteoporosis can have a harsh negative impact on the developing lineage pathway [14–22]. Tissue engineering bone for the healthy 18-year-old male patient will require a different format than for the 75-year-old male osteoporotic patient. The 18-year-old will have more osteoblasts, the osteoblasts will be more active, and the pre-osteoblasts will be more plentiful. Therefore, engineering bone for the elderly, osteoporotic individual must incorporate a factor or factors to compensate for osteoblast deficiency and dysfunction. Moreover, the elderly recipient of a tissue engineered bone construct may require cells that will differentiate into osteoblasts. There are additional physiological considerations that must be addressed for tissue engineering bone. These include gender, systemic health, and anatomic location of the engineered construct. Anatomic location includes biofunctional load and vascularity. For example, a distal tibia will have less blood supply than the proximal mandible. Bone is a vascular-dependent tissue and the challenges to support vascularity of an implanted, engineered bone tissue are more stringent in a vascular-poor region versus a region of rich blood flow, such as for facial structures. Furthermore, functional stimulation is associated with additional engineering challenges. The parietal bone will be less stringently challenged functionally than the distal femur. Therefore, structural design properties for tissue engineered flat bones of the face will be unique from the design properties engineered into femurs, for example. The purpose for the emphasis on physiological fundamentals in bone tissue engineering is to highlight the complexity and diversity of bone. However, physiological fundamentals also can provide a pathway for simplification. A specific problem can be addressed when we appreciate the fundamentals.

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In summary, the physiological considerations for bone tissue engineering must consider gender, age, anatomic location (including vascularity and function), and systemic patient profile (e.g., diabetes mellitus, steroid therapy). The rest of this chapter will deal with a basic strategy to design and develop tissue engineered bone. The strategy will include a discussion on cells, signaling molecules, and biomimetic matrices. II FUNDAMENTAL CELLS OF THE SKELETAL SYSTEM IMPORTANT FOR TISSUE ENGINEERED BONE There are many cell phenotypes associated with bone. We will focus on three: osteoblasts, osteocytes, and osteoclasts. A Osteoblasts Osteoblasts are derived from a mesenchymal cell lineage [23,24]. The progression from undifferentiated mesenchymal stem cells to end stage osteoblasts may take two routes, one leading to determined osteoprogenitor cells and the other to inducible osteoprogenitor cells [25]. The determined lineage is associated with cell condensations and may be responsible for bone formation during embryogenesis, whereas during fracture repair an inducible population may be susceptible to soluble inductive morphogens: polypeptides that promote expression of distinctive cell phenotypes, the effect being dose related [26]. Experimental studies trace the osteoblast pedigree to bone marrow stromal cells and blood vessel pericytes. The former have adipogenic and chondrogenic potential, based on quantitative availability of the morphogen bone morphogenetic protein (BMP) [27,28]. Molecular lures for pericytes, pre-osteoblast cells from endosteum, periosteum, and marrow may include fragments of collagen, as well as osteocalcin, bone sialoprotein, and gamma-carboxyglutamic acid [29]. Bone marrow stromal cells can undergo asymmetric division, with one daughter cell retaining progenitor capability and the other can differentiate to end state cells. Evidence suggests fibroblast growth factors (FGFs) and platelet-derived growth factor (PDGF) can stimulate mitogenesis in vitro and therefore may be important signaling molecules for tissue engineered bone. The process of cell recruitment during bone repair is followed by mitogenesis and, finally, differentiation. Bone morphogenetic proteins appear to be key for osteoblast differentiation [30–33]. During bone formation, osteoblasts assemble into tessellated cohort to deposit a layer of osteoid vectorially at a rate of around 1–2 m per day. After about 10 days, the osteoid calcifies at a rate of approximately 2–3 m per day. The process is finetuned through the secretion of proteins, such as type I collagen and alkaline phosphatase. As we age, osteoblast precursors and osteoblasts decrease in quantity and activity, predisposing to bone mass loss and the disease osteoporosis. B Osteocytes Osteoblasts captured by their calcifying extracellular matrix become osteocytes and are organized into interconnecting communities within bone through a gossamer of cytoplasmic networks traversing canaliculi. Cytoplasmic tendrils within canaliculi are coupled through tight junctions and function as a biomechanical and biological sensing syncytium acutely sensitive to changes in bone. As “bone sensors,” osteocytes serve to maintain bone mass. Osteocytes hover concentrically around a central canal (the Haversian canal) containing a blood vessel, nerve, and lymphatic. There are between eight and ten concentric swirls of

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osteocytes orbiting the Haversian canal. Underscoring the vascular dependency of bone cells is the fact that neither osteocytes nor osteoblasts are more than 300 m from a blood vessel. The life expectancy for osteocytes is many years, perhaps even decades [34]. Osteocytes are end stage cells incapable of renewal. Consequently, population replenishment is accomplished by osteoblast precursors. C Osteoclasts Granulocytic macrophage precursors found in bone marrow enter the circulation as monocytes and through asynchronous fusion produce a multinucleated cell up to 100 m in diameter with an average of 10–12 nuclei, known as an osteoclast [35]. In distinction to macrophage polykaryons, osteoclasts have a ruffled border, possess calcitonin receptors, produce tartrate-resistant acid phosphatase, and lack the Fc and C3 receptors [36]. Osteoclasts are relatively short-lived cells (about 2 weeks) and therefore must be renewed. Recruitment and development involves monocyte precursors, osteoblasts, and several signaling molecules. Interleukins-1, -3, -6, and -11, and probably tumor necrosis factor (TNF) alpha, along with transforming growth factor alpha expressed by osteoblasts appear to be important modulators for osteoclast development [12,37]. However, interleukin-11 may be the major controlling factor [38]. In addition, data indicate monocytes become enticed to the healing wound by fragments of fibronectin (a ubiquitous attachment factor) and extracellular matrix degradation products. Moreover, local pre-osteoblasts secrete TRANCE (a member of the TNF family), which activates pre-osteoclast-like cells through the receptor RANK [29,32]. Furthermore, macrophages at the wound vicinity express FGFs and vascular endothelial growth factors (VEGFs), prompting blood permeability and neo-angiogenesis, thereby providing transit conduits for additional monocytes to replenish those lost to injury. Morphologically, a multinucleated giant cell must be attached to bone and display a ruffled border to qualify as an osteoclast. In this format the physiological role of osteoclast-mediated bone resorption is accomplished [39]. Through as yet mysterious signals, osteoclasts are directed to cease resorbing bone, at which time they undergo apoptosis after about 2 weeks of activity. A communicative reciprocal interaction between osteoblasts and osteoclasts energizes the cellular dynamics of resorption. When osteoblasts disperse from bony surfaces in response to parathyroid hormone [40], an exposed osteoid-mineralized zone provides osteoclasts an opportunity to attach. Attachment involves surface adhesion molecules (i.e., integrins) and proteins. Osteopontin, a sialophosphoprotein, secures osteoclast docking to bone through an arginine-glycine-aspartic acid motif [41]. III FUNDAMENTAL SIGNALING MOLECULES IMPORTANT FOR TISSUE ENGINEERED BONE The human organism represents the sum integration of trillions of individual cells that replicate, grow, differentiate, perform specific functions, and die. In addition to the surrounding and interacting physical environment, from specific extracellular molecules to complex architecture, other regulatory cues are required. This continuing process from conception to death requires intercellular communication between neighboring cells in any given tissue to communication between cells in diverse and distant organs throughout the body. This continuum is enabled by cellularly synthesized chemical agents. These agents are globally referred to as hormones and include an autocrine (the cell type that secretes the hormone is directly impacted by it), paracrine (hormones produced by a given cell directly

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impacts neighboring cell types within a given tissue), juxtacrine (hormones bound to the surface membrane of one cell, while remaining bound, directly impact an immediately adjacent cell), and endocrine (hormones secreted by a given cell are transported via the circulation to distant cells in other organ systems) characterization. In addition to these local and systemic effects, hormones act via specific temporal and spatial relationships and interact with other hormones. Varying responsiveness of target cells, due to the level of target cell differentiation and other modifying factors, provides the profuse range of chemical signaling possible. Developing and understanding these processes in regard to normal and varying physiological states of the skeletal system and applying this knowledge is essential in developing more effective engineered biomaterials for bone therapies. As with all other organ systems, the skeletal system is shaped by hormones. In postnatal life, such hormones as growth hormone, insulin-like growth factor (IGF-I), parathyroid hormone, calcitonin, thyroxin, insulin, glucocorticoids, vitamin D, estrogen, and testosterone act on a systemic level to directly and indirectly modify bone growth, differentiation, function, and repair. This is exemplified by classic endocrine pathological states resulting from a breakdown in the communication pathway of a particular hormone, which leads to a loss of endocrine homeostasis and altered skeletal states. Diabetes, chronic glucocorticoid use, menopause, and acromegaly are examples of such endocrine pathologies that result in disruption of normal growth, differentiation, function, and repair processes. The reader is referred to representative texts and review articles detailing these pathophysiological processes [4,42–49]. The aging condition can also be thought of as being in endocrine imbalance as it is associated with reduced levels of a range of hormones including growth hormone, estrogen, testosterone, and IGF-I [50–57]. These conditions can be expected to adversely impact the successful outcome of any bone engineering therapy, thus their impact must be understood and alternative strategies taken to maximize outcome. Strategies can be summarized into modifying the detrimental effects of a particular endocrine pathology by affecting that pathway at a systemic level or focusing on a specific local region of treatment. Hormonal therapeutic strategies can be summarized into the positive modification of the detrimental effects of a particular endocrine pathology by affecting it at a systemic level or focusing on a specific local region of treatment. In regard to systemic delivery, aside from financial considerations and the logicistical complications of administering systemic treatments, undesired effects are likely due to the diverse and multiple target organs impacted by most hormones. The complexity surrounding systemic therapies is exemplified by Fig. 1, which presents a schematic of the simplified control pathways controlling circulating concentrations of growth hormone/IGF-I in respect to the skeletal system. Effective dosing, targeting the desired target tissue/cell type, and avoiding undesired side effects remain daunting challenges in systemic therapies, especially with hormones that can differentially impact an extensive range of tissues and can be associated with neoplastic development as well. An extensive subset of hormones includes polypeptide-based growth factors, cytokines associated with immune and hematological systems (interleukins, tumor necrosis factor, colony-stimulating factors), and prostaglandins. Although these hormones have both systemic and local effects, overall the focus for these hormones has been at the local bone level. It is beyond the scope of this review to detail the physiology of each of these hormones, thus the reader is directed to the extensive literature base, especially in regard to the broadly recognized growth factors critical for bone function, including the BMP, FGF, IGF, epidermal growth factor (EGF), platelet-derived growth factor (PDGF), transforming growth factor–beta (TGF-) families. In addition to extensive basic research implicating

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Figure 1 Simplified control system schematic of circulating plasma IGF-I regulation. The system of negative () and positive () feedback loops demonstrates the multiple levels of feedback complexity. Extensive modification of this control system results from in other interacting tissues, hormones, IGF binding proteins, and varying physiological states.

the importance of these growth factors in normal bone growth and developmental function is their recognized potential in enabling bone repair and regeneration therapies. Potential target areas of cell regulation include replication, recruitment, and differentiative function. Additionally, such growth factors impacting angiogenesis as FGFs and VEGFs offer additional target areas to facilitate bone repair [58–60]. Within the context of utilizing the various scaffold materials as local delivery vehicles for growth factors, there are a number of basic physiological concepts to be considered. First is the paracellular environment of the target local bone environment. This environment consists of four interacting constituents; cells, extracellular matrix, interstitial fluid, and growth factors. Cells, as the primary response elements, require regulation from growth factors transversing some aspect of the interstium whether of local or systemic source. Interstitial transport from the vasculature, local cell sources, or exogenous delivery involves fluid transport, the interaction with solid-phase components of the extracellular matrix, interaction with fluid-phase components of the interstitial fluid, additional sources of growth factors, target and nontarget cell metabolism, lymphatic and vascular sinks, and proteolytic processing. A schematic diagram is presented in Fig. 2 to illustrate some of these interactions in respect to providing a concentration gradient of a growth factor to a target cell population. Essential to the selection of a specific growth factor(s) is to understand the various avenues in which growth factors are made available to the bone injury site (whether the initial injury, fracture, or surgical manipulation). In the initial wound a hematoma is formed

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that, in addition to forming a blood clot to minimize blood loss, presents the first provisional extracellular matrix filling any spatial voids from either the direct injury or the placement of a scaffold into the wound site. This matrix is critical to provide an avenue for the initial cellular invasion that begins the repair process. In conjunction with this are the provision of a range of growth factors via the plasma and cellular blood component materials (i.e., platelets). The IGF-I from plasma becomes entrapped and both the serum phase of the clot as well as directly immobilized to fibrin fibrils via IGF binding proteins (IGFBPs) such as IGFBP-3 [61]. Platelets degranulate during clot formation providing a rich source of PDGF, FGF, TGF, and IGF-I as well extracellular molecules and proteases critical for repair processes. As macrophages begin to invade the clot they in turn secrete such growth factors as PDGF and IGF-I, further promoting surrounding resident cells in the bone and surrounding tissues to invade the clot. Surrounding soft tissues, primarily muscle, represent an additional source of growth factors to the injury site. Last is the resident storehouse of growth factors inherent in the organic and inorganic bone matrix. Insulin-like growth factors I and II represent the highest concentration, with much lower but significant concentrations of FGF, PDGF, TGF-, and BMPs [62–67]. Figure 3 illustrates the various endogenous sources of growth factors to healing bone. Similarly to the variation in bone quantity and quality across physiological and pathophysiological states, there is a wide variation in hormonal status and bioactivity. Emerging evidence, both basic and clinical, implies a connection between these observations and variation of bone repair. Therefore, it is likely that in regard to the inclusion of hormones as signaling agents in engineering bone therapies, that neither a single hormone nor temporal or spatial presentation regimen will satisfy all clinical situations. The individual differences between patients will surely modify a given therapy. Systemic hormonal concentrations and locally available hormones including growth factors vary greatly within and between various physiological and pathophysiological conditions. Variations in hormonal status can be associated with gender, age, anatomic location, and pathophysiological state and are likely involved with the clinical differences in repair potential between these conditions. By example, circulating growth hormone and IGF-I steadily vary with age peaking during adolescence and rapid growth normalizing until the third decade then steadily declining by as much as 10% per decade; postmenopause, estrogen circulates at minimal levels compared to premenopause; diabetes and chronic glucocorticoid usage is associated with depressed circulating IGF-I; and enhancement of hormonal systemic concentrations has potential to augment bone repair.

Figure 2 Modification of a theoretical growth factor gradient by interstitial interactions. (A) Theoretical free diffusion at an arbitrary time point; (B) diffusion based on extensive target cell metabolism of growth factor; (C) reverse diffusion gradient from cells; (D) diffusion based on growth factor interaction with delivery scaffold, endogenous fibrin clot, and/or relevant binding proteins or other soluble molecules.

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Mechanical stimulation of bone—as with other musculoskeletal tissues providing support, enabling movement, and manipulation of the environment—is critical for its normal function. Recent studies suggest interrelationships between mechanical force, hormones, and bone viability. Exercise is associated with systemic changes in a variety of hormones. At a local bone level, growth factors are likely to play a major role. These force/hormone/bone interrelationships become clinically relevant especially when the target patient is bed-ridden. The mechanical vitality of surrounding soft musculoskeletal tissues will directly impact bone repair as well [68–71]. Variation at the local bone level as a product of secretion by cells involved in the repair processes and growth factors mobilized from available bone matrix stores are also likely to modify bone repair. Local bone matrix concentrations will vary with age, anatomical site, and under pathophysiological conditions [72–74]. This is depicted in Fig. 4. Finally, understanding and utilizing the specific physiology of a growth factor targeted for therapy is crucial to the design stream. Essentially every growth factor is secreted as a mature protein, however, once in the interstitial space a wide range of biological processes modify its potential bioactivity. Transforming growth factor–beta is secreted bound to its cleaved preprotein fragment [75,76]; IGFs, TGF-, and BMPs are rapidly sequestered by specific binding proteins that modify receptor bioavailability. Through these binding proteins as well as direct growth factor binding to extracellular components [61,75–79] creates the matrixbound store of growth factors including IGFs, TGF-, BMPs, PDGF, and FGFs. These various binding interactions determine the availability of a given growth factor, influencing receptor binding; presentation of a growth factor as a soluble protein-, cell-, or matrix-bound growth factor, transport; and residence of a growth factor within a tissue site. These same interactions are also expected to control the spatial and temporal availability of growth factors. Strategies based on these concepts will allow a more biomimetic application. Interaction of IGF-I with hydroxyapatite is but one example: IGF-I binds only nonspecifically to hydroxyapatite or fibrin, but in the presence of IGFBPs it becomes sequestered [61,79]. Strategies including IGF-I alone would be expected to result in an almost instantaneous re-

Figure 3 Schematic of the systemic and local tissue sources of representative growth factors to the bone wound site.

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Figure 4 Variability of IGF-I to the bone wound site as a function of physiological status. Downward arrows indicate a reduced IGF-I concentration. Question marks indicate that the effect is unclear. lease of IGF-I from a hydroxyapatite or fibrin surfaces to the interstitial fluid. Thus any sustained release results from the structural complexity of the scaffold and/or the infiltration of the additional blood clot with included IGFBPs from plasma and platelet components to regulate IGF-I binding to scaffold. These concepts as well as the cell level control systems, including negative feedback mechanisms and interactions between growth factor systems, must be appreciated or the development of hormonal therapies in bone repair will consist of an Edisonian approach of trial and error. Considering the astronomical number of possible permeations, the expense of basic and clinical research, and the time lines involved, the selection of a growth factor or combination of growth factors must be based firmly on available and developing physiological concepts. This is universal irregardless of the scaffold, the anatomical site, the patient health status, and most especially the delivery methodology of growth factor—systemic/local, protein in solution, bound to scaffold, released from bioresorbable scaffold components, or by a genetically manipulated cell. IV BIOMIMETIC MATRICES IMPORTANT FOR TISSUE ENGINEERED BONE Nature has provided us with the perfect template for the design of biomaterials. Beyond the use of polymers or ceramics, one must consider additional factors that will improve and hasten healing. When there is an injury in the body, there is a specific physiological cycle that is followed. An understanding of that cycle must be known. As discussed in this chapter, growth factors and various extracellular (ECM) proteins, which regulate cell attachment, reside in and are therefore an inherent component of native bone. These proteins provide the osteogenic capacity observed with autograft and allograft. That is, these components provide the chemical cues to initiate and maintain those processes associated with bone formation. Synthetic bone graft materials, whether polymer, ceramic, or metal, provide only an inherent osteoconductive capacity. In other words, they provide only a physical platform by which they support cell attachment, shape, constrain, and control tissue development. By incorporating such biological components within synthetic bone scaffolds a closer approximation of native bone graft material is achieved. In addition, control of the specific addition of scaffold and biological components minimizes immunological rejection potential. Thus a synthetic bone graft can be achieved with the inherent os-

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teogenic properties of allograft, without the immunological rejection or disease threat potential, and without the complications often associated with the graft donor site. Furthermore, by controlling growth factor concentrations, these synthetic bone grafts have the capacity to be more osteogenic then either autograft or the receiving host bone bed. Such tissues are often depleted of endogenous growth factor stores because of various disease and physiological states, including diabetes, osteoporosis, anti-inflammatory drug therapy, and old age. The further addition of cellular components to synthetic scaffolds provides yet another step toward achieving a truly pseudonative bone graft. Although actual surviving cell numbers may be rather small, these cells provide for colonization focal points within the bone graft. This improves the incorporation potential of the graft with the surrounding bone, which would otherwise be limited to cellular colonization from the surrounding host bone bed. By controlling the cell types utilized in these processes, there is potential to improve rates of osteointegration, the incorporation of host bony tissues throughout the graft. This is especially significant in various disease and physiological conditions where available host cell populations may be compromised. Biomaterials that support the incorporation of cells and signaling molecules are being widely studied [80]. Synthetic biomaterials are typically comprised of biodegradable polymers, such as poly(D,L-lactic-co-glycolic acid). Native matrices include collagen and hyaluronic acid. Numerous methods have been developed to create three-dimensional porous scaffolds for tissue engineering. These techniques include particulate-leaching [59,60], emulsion/freeze-drying [83], temperature-induced phase separation [84–87], gel casting [66,67], gas foaming/particulate-leaching [90,91], and fiber binding meshes [92]. These different procedures result in varying scaffold microstructures. For example, the emulsion/freeze-drying method can result in pore sizes ~35 m. However, this process has proven to be difficult to reproduce. Particulate-leaching creates larger pore size ranges by incorporating sieved sodium chloride or soluble sugar particles of a desired size. Similarly, the temperature-induced phase separation method offers a controlled approach to fabricate low density polymer scaffold structures exhibiting pore sizes in the 50–400 m range. Thermally induced gelation combined with solvent exchange and freeze-drying creates a nanoscale fibrous extracellular matrix [93,94]. Current methods of signaling molecule incorporation include adsorption of growth factors to the surface prior to implantation [95] and incorporation of growth factors during the scaffold fabrication process [96]. Babensee et al. has written a comprehensive review on growth factor delivery [97]. Additionally, microparticles containing growth factor can be injected into a porous scaffold [98]. Growth factors can be encapsulated in microparticles (mean diameter 1–1000 m) or nanoparticles (mean diameter 1–1000 nm). Drugs and other growth factors have been incorporated into PLGA microparticles, such as human growth hormone [99], Japanese encephalitis virus vaccine [100], vascular endothelial growth factor [101], insulin-like growth factor [102], transforming growth factor–1 [50,51], fibroblast growth factor [105], cyclosporine [106], and cisplatin [107]. Growth factor incorporation into polymer nanoparticles has been less widely studied [108–111]. When signaling molecules are incorporated into biomimetic matrices, mild operational conditions are employed to avoid denaturation of these molecules. For example, the emulsion/freeze-drying method allows for the incorporation of peptides, but is not used for the incorporation of DNA due to high shear field forces generated by the homogenizer. Other methods that require heat or toxic solvents are also unsuitable. Thus, most available

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fabrication techniques are excluded by these limitations, while others require modifications for the incorporation of macromolecular agents. Particulate-leaching is a scaffold fabrication technique that can result in a porous structure with incorporated signaling molecules. However, this technique yields an uneven distribution of scaffold density throughout the matrix. When the mixture of salt particles and polymer solution is cast into a mold, salt particles often sink to the bottom of the mold. After solvent evaporation and porogen leaching, the scaffold consists of a denser layer at the polymer/air interface. Typically, that dense polymer layer is removed by milling or slicing. However, if signaling molecules are incorporated within that scaffold, the removal of a portion of the scaffold would result in a loss of signaling molecule. To avoid the thick polymer surface, Agrawal et al. has added vibration to the fabrication process. During solvent evaporation, vibration maintains an even dispersion of salt particles throughout matrix, resulting in a homogeneously porous matrix [112]. While continuing to use traditional delivery systems for signaling molecules, such as microspheres [50,51], hydrogels and polymer reservoirs [75,76], and capsules [115], investigators are making efforts to incorporate these delivery systems directly into biomimetic matrices. These delivery systems can be incorporated into the biomimetic matrix either during or after scaffold fabrication. Examples include localizing growth factors in a porous polymer scaffold prior to implantation [95] and infiltrating microparticles containing growth factors into the fabricated porous scaffold [98]. Mooney et al. has incorporated both growth factors and DNA during the scaffold fabrication process by a high-pressure gas-foaming/particulate-leaching technique [58,78]. Another methodology is the incorporation of growth factors contained in polymeric microspheres into the porous polymeric scaffolds during scaffold fabrication [117]. The basic tenet of this technique is to transiently protect the microspheres with a water-soluble coating which resists the organic solvents used in the particulate-leaching method. Without this protection the microspheres will prematurely dissolve when incorporated in the larger polymeric scaffold. The watersoluble coating is removed during the particulate-leaching step to remove the porogen, thus creating the porosity in the scaffold. Modification of the microsphere polymer chemistry permits differential degradation rates, thus allowing temporal control of delivery. Loading different growth factors at varying concentrations permits control of the source concentration. The inclusion of biological control agents remains a basic tenet before viable success using biomimetic matrices can be fully achieved. Growth factors, as the principal extracellular biomolecules regulating bone growth, development, and repair, must be included in such a manner as to induce the surrounding host bone tissues to colonize the inserted synthetic construct and ultimately replace it with viable bone. Although the optimal combinations of specific growth factors and temporal relationships are likely to be specific for a given graft placement location and the physiological status of the recipient, there are common basic precepts which need to be established that are applicable across any specific modifier. The controlled delivery of growth factors within scaffold materials will require more exacting methods than are currently available. V CONCLUSIONS The successful incorporation of any tissue engineered material within a host body is dependent on the appropriate communication between cells, tissues, and the host system as

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a whole. The growth and differentiation of cells into organized tissues and tissues into an integrated functional organism require chemical communication across this entire spectrum of hierarchies, including cell-to-cell, cell-to-tissue, tissue-to-organism, and cell-toorganism interactions. These communication interactions are not only absolutely essential for an organism’s initial development, but are also essential in the maintenance of an organism’s functionality throughout that organism’s lifetime. Of specific interest to tissue engineering applications is the role of chemical communication during tissue repair and regeneration as regulated by specific protein hormones referred to as growth factors. Investing and maintaining biofunctionality will involve research on the interaction of cells and signaling molecules within the scaffold matrix to optimize fusion of layers, bone growth potential, and integration into surrounding host tissues. Direct and selective placement of cell populations throughout the scaffolding will provide the means to systematically study and control cellular material incorporation. Signaling molecules affecting the rate and limitations on cell growth in scaffolds and the state of differentiation and responsiveness of the resulting expanded populations provide further enhancement and control of tissue growth. The understanding and successful manipulation of cells and growth factors in growth, development, repair, and regeneration of tissues are thus major linchpins of tissue engineering. REFERENCES 1. Boskey A. 1994. Bone and cartilage mineralization. In: Bone Formation and Repair. Brighton C. T., Friedlaender G., Lane J. M., Eds. American Academy of Orthopaedic Surgeons: Rosemont, pp. 23–38. 2. Buckwalter J. A., Glimcher M. J., Cooper R. R., Recker, R. 1995. Bone biology. Part II: formation, modeling, remodeling, and regulation of cell function. J. Bone Joint Surg. 77A(8):1276–1289. 3. Buckwalter J. A., Glimcher M. J., Cooper R. R., Recker R. 1995. Bone biology. Part I: structure, blood supply, cells, matrix, and mineralization. J. Bone Joint Surg. 77-A(8):1256–1275. 4. Bilezikian J. P., Raisz L. G., Rodan G. A., Eds. 1996. Principles of Bone Biology. Vol. 1. Academic Press: New York, pp. 1–1398. 5. Einhorn T. 1996. Biomechanics of bone. In: Principles of Bone Biology. Bilezikian J., Raisz L., Rodan G., Eds. Academic Press: San Diego, pp. 25–37. 6. Hollinger J. O., Buck D. C., Bruder S. 1998. Biology of bone healing: its impact on clinical therapy. In: Tissue Engineering: Applications in Maxillofacial Surgery and Periodontics. Lynch S., Marx R., Genco R., Eds. Quintessence: San Diego, pp. 17–53. 7. Baron R. 1994. The cellular basis of bone resorption: cell biology of the osteoclast. In: Bone Formation and Repair. Brighton C. T., Friedlaender G., Lane J. M., Eds. American Academy of Orthopaedic Surgeons: Rosemont, pp. 247–252. 8. Bruder S. P., Fink D. J., Caplan A. I. 1994. Mesenchymal stem cells in bone development, bone repair, and skeletal regeneration therapy. J. Cell Biochem. 56:283–294. 9. Bruder S. P., Kurth A. A., Shea M., Hayes W. C., Jaiswal N., Kadiyala S. 1998. Bone regeneration by implantation of purified, culture-expanded human mesenchymal stem cells. J. Orthop. Res. 16:155–162. 10. Aubin J. E., Liu F., Malaval L., Gupta A. K. 1995. Osteoblast and chondroblast differentiation. Bone 17:77S–83S. 11. Caplan A. I., Bruder S. P. 1997. Cell and molecular engineering of bone regeneration. In: Principles of Tissue Engineering. Lanza R., Langer R., Chick W., Eds. Academic Press: San Diego, pp. 603–618.

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3 Mimicking the Natural Tissue Environment Christopher J. Woolverton Kent State University, Kent, Ohio Judith A. Fulton Akron General Medical Center, Akron, Ohio Stephanie T. Lopina University of Akron, Akron, Ohio William J. Landis Northeastern Ohio Universities College of Medicine, Rootstown, Ohio

I INTRODUCTION Tissue engineering essentially involves the application of molecular and cellular biology to resolve pathologies and augment natural processes in order to assist with tissue development, repair, and healing. While the initial focus was on successful use of implants, the concept of tissue engineering today encompasses the integration of natural and synthetic materials to mimic normal tissue structure and function. Tissue engineering is the uniting of many discrete areas of knowledge found in scientific, technological, and medical disciplines in an effort to understand and design functionally equivalent tissues. Design and synthesis of biomaterials, analysis of biocompatibility and biostability, as well as dissection of the normal physiological processes that result in tissue formation and homeostasis are all components of modern tissue engineering. Furthermore, tissue engineering begs the combining of appropriate ingredients (e.g., cells, matrix, cytokines, soluble proteins, low molecular weight solutes) in the correct environment in order to create new tissue. What are the proper ingredients? How much of each is needed? When should each be supplied to the process? These are valid questions one needs to answer when considering the engi43

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neering of a tissue. Clues and answers can certainly be obtained from the natural course of events in the wound healing process. The lofty goal of this field is to repair defects and resolve pathologies that would otherwise not occur by homeostatic mechanisms or surgical intervention. Thus, a thorough understanding of the natural materials and processes is required for tissue engineering to succeed. The main focus of the body’s response to injury is repair and homeostasis for the purpose of survival. Therefore, the first response after injury is hemostasis, which is accomplished by vasoconstriction and blood clot formation. Next, efforts are made to combat infection and eliminate damaged tissue and foreign material. At the same time, the body must replace damaged tissue, supply nutrients, remove wastes, and cover exposed surfaces to prevent desiccation and microorganism intrusion. Finally, the new fragile tissue must be strengthened to prevent subsequent breakdown. All this is effected through a series of complex, coordinated, and sequential events. The wound healing process is typically divided into three overlapping phases: inflammation, tissue formation, and remodeling. During the inflammatory phase, neutrophils and macrophages enter the wound site in response to chemotactic factors created by damaged tissue and released by platelets. The inflammatory cells phagocytize bacteria and debris, thus providing natural antibiotics and debridement. In addition, macrophages secrete numerous bioactive molecules that help to orchestrate the ensuing tissue formation phase. This phase is highlighted by migration to the wound of specialized cells whose function is to replace lost tissue. Fibroblasts synthesize and secrete new matrix material, keratinocytes create a new epithelium, and endothelial cells establish a vascular network. Finally, the newly formed matrix is modified and remodeled to create strength and durability. The end result is new tissue that enhances survival but is not necessarily identical to the original lost tissue. Function may yet be impaired, and extensive injuries require medical interventions such as reconstruction or transplantation. Excellent reviews and detailed discussions on wound healing can be found elsewhere [1–3]. Although cells are the key players in the formation of new tissue, they cannot function without the appropriate extracellular matrix (ECM). The matrix influences the cells and the cells in turn modify the matrix, thus creating a constantly changing microenvironment throughout the healing process. The ECM begins with the blood clot and ends as scar tissue, transformed along the way in an orderly and sequential fashion to meet the needs of the healing wound. Matrix development has been divided into three main phases. The first provisional matrix is the fibrin clot; the second provisional matrix is composed primarily of fibronectin and hyaluronic acid; and the collagenous scar provides the final, more permanent matrix [4]. However, the wound healing ECM may more appropriately be considered a matrix continuum that gradually changes and moves through each of the three provisional matrix phases. A Chapter Organization This chapter presents the known literature defining natural components of the tissue microenvironment so that insightful recapitulation of the natural process can be considered during the engineering of tissue ex vivo. A discussion of criteria used to mimic the natural tissue components follows so as to explain the scientific rationale for the development and selection of biomaterials currently applied in tissue engineering. While not an exhaustive list, the principal materials now used and included in the chapter are fibrin, hyaluronic acid, collagen, dermal explants, intestinal submucosa, alginate, and chitin. Synthetic materials

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that play a pivotal role in tissue engineering are also described to provide the reader with an appreciation for their particular utilization in the field. Finally, two specific clinical applications are detailed to give examples of particular cases in which tissues are engineered using the theory and techniques described herein. II THE NATURAL TISSUE ENVIRONMENT Natural tissue development of a wound (repair, remodeling, and so on) typically begins with a blood (fibrin) clot which adheres to surrounding tissues and fills the wound voids. The process of clotting is designed to rapidly stop blood loss caused by injury. This process occurs by means of three distinct events: aggregation of platelets to seal leakage, vasoconstriction to reduce blood flow, and formation of an insoluble protein mesh to cover the ruptured site. It is of some note that while platelets act as a physical barrier to initially seal the vascular breach, they also provide a source of cytokines, growth factors, and other low molecular weight molecules required for natural tissue repair and remodeling. The signal that initiates platelet aggregation is the exposed collagen fibers of the endothelium. Platelets adhere to the exposed collagen and subsequently change shape while releasing heparin-binding protein, phospholipids, thromboxane A2, serotonin, ADP, factor V, and other molecules. The exposed collagen also recruits circulating von Willebrand factor (and factor VIII) that facilitates platelet sealing of the endothelium [5]. In addition to recruiting platelets, the damaged endothelium initiates the blood-clotting pathway, and exposed extravascular membrane proteins serve as receptors for soluble zymogens and cofactors in the plasma. The clotting process results from the activation of a series of zymogens in a cascade fashion with the subsequent formation of fibrin. In the final steps, blood coagulates through the action of thrombin, a proteolytic enzyme, on fibrinogen, a soluble protein normally found in circulating blood at a concentration of 3.0 mg/mL. Thrombin cleaves fibrinogen molecules to form fibrin monomers that associate into fibrils and eventually assemble into a three-dimensional fiber network. As the fibrin polymerizes, other proteins and blood constituents are bound to or trapped within the forming matrix. The most prominent protein species is fibronectin, which normally circulates in the blood at a concentration of 0.3 mg/mL. Other components include vitronectin, von Willebrand factor, thrombospondin, tenascin, denatured collagen, and cytokines. Crosslinking between fibrin molecules and between fibrin and other ECM molecules is catalyzed by activated factor XIII (transglutaminase). Platelets also interact with the fibrin clot, adding additional fibrinogen and growth factors. This first provisional matrix adheres tightly to the local wound site of the tissue and provides a scaffold for entry and migration of the cells necessary for the healing process [4,6]. Although most cells entering the wound site contribute to the ever-changing matrix, fibroblasts are credited with providing the majority of the cellularly secreted ECM molecules. Early in the repair/healing process, fibroblasts secrete fibronectin and hyaluronic acid, a glycosaminoglycan (GAG). Thus, the second provisional matrix is laid down while the fibrin-rich matrix is digested through the action of proteases. In the same manner, as healing and repair progress, the fibronectin/hyaluronic acid matrix is gradually degraded and replaced with a collagen matrix. Fibroblasts initially deposit type III and then type I collagen, utilizing the fibronectin matrix as a structural blueprint [4]. Type I and type III collagens are fibrillar proteins with a triple helical conformation. In order to provide strength to a healing wound, collagen fibrils steadily aggregate to form fibrous bundles. At the same time, fibroblasts deposit proteoglycans, which are composed

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of sulfated GAGs (for example, chondroitin-4-sulfate and dermatan sulfate) with protein cores. When a sufficient amount of collagen has been deposited, fibroblasts cease production of new collagen and begin the process of collagen reorganization. The phenotype of fibroblasts changes to that of myofibroblasts, assisting in contraction of the wound, compaction of the matrix, and scar formation. Finally, as the scar matures, collagen is degraded and resynthesized, thus allowing for remodeling through the formation of thicker bundles and crosslinking [4,7,8]. The ECM materials used by the human host are ideally suited for their purpose and create a “cell-friendly” environment. For example, integrin receptors on migrating cells recognize many of the matrix molecules such as fibrin, fibronectin, vitronectin, and collagen, and the cells presumably move through the matrix by means of contact guidance. On the other hand, the high levels of hyaluronic acid, a hydrated polysaccharide, permit easier movement of cells through the matrix by virtue of their low impedance properties. In contrast, the stiffer collagen matrix adds strength to the wound at a time when cellular migration is no longer necessary. The combination of collagen and proteoglycans is also ideal because the triple helical fibrils of type I collagen aggregate into thick bundles that increase tensile strength while the proteoglycans, with a protein core and flexible GAG side chains, add resilience. In addition, as various cell types migrate into the matrix, their new environment stimulates them to secrete a variety of nonmatrix, bioactive molecules (such as proteases, inhibitors, and cytokines), and the ECM becomes a reservoir of these components. For example, growth factors within the matrix are protected from degradation by proteases present in the surrounding environment, and the rate of matrix degradation may be limited by bound protease inhibitors [4]. Many of the ECM materials just described have been used with some success as tissue engineering matrices. However, in the natural wound environment, these matrix molecules are not found individually in pure states, but are intermingled with a variety of other matrix and bioactive molecules. In fact, when cells are seeded on pure matrices containing only a single molecular species, they generally do not spread, migrate, or proliferate as readily as when matrices mimic the natural environment containing many different molecules. As additional knowledge is gained of the interactions between cells, matrices, and other bioactive molecules, tissue engineering matrix combinations will be discovered or synthesized that more closely mimic the natural healing process. Eventually we should move beyond the process of repair for the purpose of survival and step into a new realm of cell, tissue, and organ regeneration. III SELECTION ISSUES The localized microenvironment in which a tissue develops must support structural as well as temporal changes to facilitate the formation of final cellular organization. This is mediated by specific types and concentrations of ECM molecules that appear at defined times to direct tissue development, repair, and healing. The ECM components are constantly remodeled, degraded, and resynthesized locally to provide the appropriate type and concentration of protein with respect to time. The ECM is composed of three distinct components: insoluble protein fibers such as fibrin or collagen, soluble multiadhesive matrix proteins, and highly viscous proteoglycans. Collagen is the major insoluble fibrous protein found in the ECM. Secreted by fibroblasts and epithelial cells, collagen serves the primary purpose of withstanding stretching. Multiadhesive matrix proteins such as fibronectin, vitronectin,

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and thrombospondin are long flexible molecules that bind various collagen types, polysaccharides, growth factors, hormones, cell-adhesion molecules, and other matrix proteins. Their primary function is to link cells to the ECM and support tissue formation. Proteoglycans are a diverse group of macromolecules comprising a core protein with one or more attached polysaccharide chains [9]. Various combinations of these ECM products are required to direct the biosynthesis of each new tissue type. While the ECM can be thought of as the scaffolding in which cells are housed during their association, the ECM is by no means inert. Cellular receptors for and specific behavioral responses to ECM molecules suggest that the ECM communicates directly and/or indirectly with cellular signaling pathways to elicit specific biological functions [9]. Thus, guided orientation, mitosis regulation, and cellular differentiation appear to be directed by cellular interaction with ECM components. Selection of ECM equivalents during tissue engineering should likewise mimic these natural processes to the maximal extent possible. When selecting molecules for tissue engineering, one should understand the temporally orchestrated biological function of each species chosen. Two issues require thought in this context; biocompatibility and biostability. For the former, a thorough knowledge of an ECM equivalent should include its ability to fulfill any structural or informational responsibility. Will the material be needed early to assist in stabilizing and bridging adjacent tissues? Are cells required to penetrate and populate the material? Are specific chemical side groups required for cell adhesion? How long should the equivalent biological signals be issued, and is a greater or lesser concentration required at any given time? Besides such physiological effects, biocompatibility also includes understanding pharmacological effects, including potential toxicity of an ECM mimic. Will the material be toxic to any of the present or subsequently developed proteins, cells, or other biological factors of the engineered tissue? Will the material sensitize the host and provoke immune responses? In the case of biostability, potential ECM equivalents should mimic, as much as possible, the physical properties and “life span” of their natural counterpart materials. For example, fibrin in a wound is relatively poor at withstanding stretching because of its low tensile strength. It must be degraded as any associated collagen matrix is produced to allow for tissue flexibility without tearing (higher tensile strength). Therefore, a fibrin substitute should be more susceptible to endogenous proteases (limiting its “life span”) compared with a collagen mimic, which should not be as readily degraded. In summary, selection of appropriate biomaterial mimetics should approximate the chemical, temporal, and structural integrity of the natural counterpart in terms of physical properties, arrival and degradation, compatibility, and stability. Understanding the natural tissue environment can offer insight into the use of natural and synthetic products to engineer new tissue, repair pathologies, and produce transplantable tissue equivalents. The ensuing discussion offers examples of biologically derived and chemically synthesized biomaterials for tissue engineering.

IV CURRENTLY USED MATRIX MATERIALS OF BIOLOGICAL ORIGIN The scope of this chapter does not permit an exhaustive review. Selected examples of some principal materials used in tissue engineering and their recent development and applications are presented herein.

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A Fibrin Fibrin, the main component of the first provisional matrix at a wound site, has historically been used as a hemostatic agent and a tissue adhesive. More recently it has served as an engineering matrix for numerous types of tissues. Fibrin is formed from fibrinogen, a 340kDa protein. The fibrinogen protein (factor I) is composed of two tripeptides of an  structure. The two tripeptide structures are held by disulfide bonds at their N-terminal ends. Three globular domains are evident in the fibrinogen molecule, one at each end and one in the middle, separated by rodlike configurations. Charge repulsions prevent fibrinogen molecules from self-assembly. However, enzymatic cleavage of the central (N-terminal)  and  regions by thrombin (factor II) facilitates self-assembly of fibrinogen proteins, forming the “soft” clot [5]. A stabilized clot forms by the subsequent action of a transglutaminase (factor XIII) covalently linking adjacent fibrin monomers. The modern use of clotting proteins from blood to mediate hemostasis and facilitate wound repair dates back to the early 1900s. In 1944 purified fibrinogen and thrombin solutions were used clinically to treat hemorrhage [10]. In 1972, fibrin sealant (FS) was developed in Vienna whereby the cryoprecipitate of plasma was mixed with a bovine thrombin solution to form a fibrin clot [11]. Today, several commercial laboratories manufacture FS components for the intended use of hemostasis and wound healing applications. FS is especially useful for treatment of large surface area wounds and for sealing microvascular leakage resulting from surgery [12]. Human FS components (fibrinogen, thrombin, and factor XIII) are isolated from fractionated human plasma and treated to neutralize and/or remove microorganisms, including enveloped viruses [13], according to FDA regulation. The fibrin formed ex vivo from commercial FS systems is specifically formulated for hemostasis applications and therefore is denser than fibrin formed naturally in vivo. Thus, in-depth structure–function studies must evaluate the relative concentrations of each component needed to form the wound-equivalent fibrin matrix so that cells and other materials mimic more readily the natural repair process. The composition of a fibrin gel, as well as the types of seeded cells, can greatly influence the success of a fibrin–cell construct. For example, keratinocytes seeded on a commercially obtained fibrin gel of high purity failed to spread or proliferate unless fibronectin was added or the gels were crosslinked with factor XIII [14]. Alternatively, an impure fibrin matrix or coculture of complementary cell types can create a more favorable cellular environment. Thus, a skin equivalent with possible use as a graft material for burn patients was successfully constructed using both epidermal and dermal cells in fibrinogen obtained from human plasma cryoprecipitates (less pure than commercially prepared fibrinogen). Human fibroblasts were seeded into the fibrin gels before coagulation, and human keratinocytes were subsequently seeded on top of the fibrin–fibroblast gels. After 15 days of culture, skin bilayers were obtained with positive staining for basement membrane proteins (type IV collagen and laminin). When the composites were grafted onto athymic mice, they developed a stratified, cornified epidermis resembling that of humans [15]. In a clinical setting, Pellegrini et al. [16] successfully treated seven burn patients with autologous keratinocytes cultured on fibrin. They determined that keratinocytes obtained from skin biopsies and cultured on fibrin in the presence of irradiated 3T3-J2 cells formed large autografts bearing stem cells. Full thickness burns were debrided and grafted with allodermis, and 2 weeks later the fibrin-cultured keratinocytes were applied. All patients showed 100% epidermal regeneration of full thickness burns after 1 month, credited to the ability of fibrin to support stem cell growth [16]. In order to reduce culture times and sim-

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plify applications, Horch et al. [17] suspended subconfluent cultured human keratinocytes in a commercially available fibrin glue and sprayed the fibrin–cell matrix onto full thickness wounds on athymic mice. The fibrin polymerized in situ and re-epithelialization was complete in 7 to 14 days. The epidermal–dermal junction zone formed at a faster rate than with wounds covered with standard cultured epithelial sheet grafts [17]. Although examples of fibrin used for dermal repair are more prevalent, other tissue types have been examined as well. Work completed by Silverman et al. [18] has demonstrated that fibrin seeded with chondrocytes will support the growth of neocartilage that can adhere to native cartilage. Swine chondrocytes at a concentration of 40 million/cc were suspended in porcine fibrin glue and injected before polymerization into the subcutaneous space on the back of athymic mice. By 6 weeks, solid homogeneous cartilage had formed with histology similar to swine articular cartilage, and the presence of type II collagen (found only in cartilage) was confirmed [18]. Further studies indicated that when the fibrin/chondrocyte mixture was sandwiched between two discs of native cartilage and implanted subcutaneously in nude mice, the resulting neocartilage adhered closely to the native cartilage. Mechanical testing, in particular, tensile strength, showed the engineered construct to be significantly superior to cartilage discs joined by fibrin glue alone [19]. Peripheral nerve regeneration possibilities in fibrin have also been explored. Chick embryo dorsal root ganglia were embedded in fibrin gels of varying concentrations, and neurite growth was measured after 24 h of culture. Gels prepared with lower concentrations of fibrinogen facilitated an increase in neurite length [20]. Neurite growth was further enhanced by modifying fibrin gels with covalently bound heparin-binding peptides [21] or neuroactive peptides such as laminin or N-cadherin [22]. Additional experiments in vivo used a rat severed nerve model in which tubes containing various test matrices were placed in the gaps between excised nerve stumps. The number of axons that grew into the tubes filled with fibrin/laminin derivatives was 85% greater than the controls with unmodified fibrin [22]. B Hyaluronic Acid Hyaluronic acid (hyaluronan) is a major component of the secondary provisional matrix and has also been used in numerous tissue engineering applications. Since hyaluronan is a water-soluble hydrogel and forms a viscous, jellylike solution, it is generally modified by esterification to make it insoluble so it can be fabricated into gels, films, or woven materials. The benzyl ester hyaluronan products HYAFF-11™ and LaserSkin™ (Fidia Advanced Biopolymers, FAB, Abano Terme, Italy) have been used to engineer skin bilayers in vitro [23]. In addition, tissue similar to hyaline cartilage was produced by human nasoseptal chondrocytes seeded on HYAFF-11 and incubated in athymic mice [24]. Another product, Vivoderm™ (ConvaTec, Princeton, NJ), was used to culture autologous keratinocytes, which were then successfully applied to chronic full thickness ulcers on three patients [25]. Benzyl ester hyaluronan has also been combined with gelatin (hydrolyzed bovine collagen) to form a composite sponge that was seeded with bone marrow–derived mesenchymal progenitor cells. Culture in the presence of transforming growth factor–1 stimulated cells to produce a type I collagen-rich ECM [26]. Subcutaneous implantation in athymic mice yielded bone and cartilage tissue [26]. Hyaluronan has alternatively been crosslinked with glutaraldehyde and then coated with polylysine. These modifications produced a matrix that promoted Schwann cell attachment and proliferation and thus has potential nerve graft application [27]. Another unique composite containing hyaluronan, a

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polyanion, and polypyrrole, a polycationic electronic conductor, provided a matrix with both biological and electrical activities. Experiments in vivo showed this composite also possesses angiogenic properties, making it a candidate for bone or nerve regeneration applications [28]. C Collagen Collagens are major components of the ECM and facilitate, integrate, and maintain the integrity of a wide variety of tissues. They serve as structural scaffolds of tissue architecture, uniting cells and other biomolecules. Many physiochemical studies have elucidated the physical and biochemical properties of collagens. As proteins, they have a generic structure based on their linear amino acid sequence. A repeating tripeptide sequence of glycineX-Y, with abundant proline in the X position and hydroxyproline at the Y position, imparts a helical secondary structure to the nascent collagen protein. This helix, however, is distinct from a traditional -helix and is termed the polyproline II helix. The effect of having glycine repeated every third residue in the polypeptide results in a unique sequence. Glycine is the only amino acid that can be packed tightly between other molecules. Thus, the linear sequence of collagen can naturally form a semirigid helix with glycine providing hydrogen binding sites to the other groups of collagen polypeptides. The biochemical effect of this type of protein architecture is the formation of a supersecondary structure resulting from the union of three polyproline II helices, stabilized by intrahelical hydrogen bonding [29]. Collagens make up a superfamily of more than 30 different polypeptides, forming at least 19 different collagen types. The triple helices of different collagens have varying flexibilities, thermal stabilities, and secondary structures, depending upon the residues occupying the X and Y positions in the repeating tripeptide sequence. Amino acid and gene sequence data have facilitated the grouping of various collagenous products. For example, fibrillar collagens (types I, II, III, V, and XI) have distinct structural features reflecting their highly conserved exon–intron gene organization. The expression of this gene organization results in fibrillar molecules consisting of uninterrupted helical domains. The fibrillar association of collagen is largely controlled by noncovalent bonds between polar and hydrophobic regions of the individual collagen strands. Thus, the lateral union of collagen strands forms fibrils spontaneously in vitro. However, there is more stringent control over fibril associations in vivo. Collagen fibril formation and the formation of polygonal protein arrangements are numerous and may represent tissue-specific structure formation resulting from additional protein modification in situ [29]. Since type I collagen is the primary matrix molecule in the final scar tissue, it has been the logical choice for many experimental and clinical applications. Historically, collagen has been prepared as gels, sponges, or composites and has been used for temporary or permanent wound dressings. An advantage of a collagen dressing is its inherent ability to attract endogenous cells and provide a scaffold for their growth and development. Soluble collagen can be readily prepared from tendon or dermis by acid extraction. The technique preserves the fibril ends and makes them available for future self-assembly and gel formation when pH is readjusted. Collagen gels have been used clinically to create skin equivalents for treating burns and excised giant nevi [30]. Autologous or allogenic fibroblasts isolated from normal skin samples were expanded in culture and seeded into bovine type I collagen gels [30]. After 4 to 6 weeks, the resultant dermal equivalent was grafted onto a wound bed followed by application of an epidermal layer 2 weeks later. By 9

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months, the quality of the healed skin, evidenced by the formation of rete ridges and deposition of elastic fibers, was superior to that observed with alternative methods [30]. A commercially available living skin bilayer composed of fibroblasts and keratinocytes in a type I collagen matrix (e.g., Apligraf™, Organogenesis, Canton, MA) has also been used to treat a variety of dermal ulcers [31–33]. In another example, skeletal muscle was prepared from a myoblast cell line and seeded in rod-shaped bovine type I collagen gels [34]. Culture resulted in thin tissue with high cellular density oriented unidirectionally. Without a vascular supply, however, tissue size was limited. When such muscle constructs were subsequently incubated in subcutaneous pockets on the backs of athymic mice, the muscle tissue grew in size. After 2 weeks, myotubes began to form, and capillary networks covered the surfaces and penetrated the interior of the engineered muscle [34]. Collagen gels have shown limited success in support of culturing corneal cells. However, type I collagen processed into porous fibrillar sponge promoted normal morphology of cultured epithelial and endothelial cells as well as keratocyte infiltration, attachment, and matrix production. In addition, keratocytes cultured in a collagen sponge matrix transmitted five times more light than collagen gel cultures [36]. Collagen sponges with varying pore sizes have also been successfully used for construction of skin bilayers [35]. Human fibroblasts have been first seeded on a highly porous collagen sponge and then layered with another sponge. The second sponge had a smaller pore size and was seeded with keratinocytes. Coculture of the two cell types allowed interaction between developing dermal and epidermal layers so that the skin equivalent was ready within 3 to 5 days for successful grafting onto athymic mice [36]. A tissue engineering matrix placed in any wound environment will be naturally degraded and replaced by a new matrix synthesized by exogenously seeded cells or by migrating endogenous cells. In this process, the rate of matrix degradation must not exceed the rate of matrix synthesis in order to prevent net loss of matrix and allow fulfillment of its function. To provide enhanced stability to collagen matrices, various crosslinking methods including UV radiation, hydrothermal treatment, and chemical methods (e.g., use of glutaraldehyde, carbodiimides, or polyamines) have been employed [37–39]. In addition to crosslinking, other modifications have often been made to impart greater biocompatibility. In these situations, collagen is often coprecipitated with the GAG chondroitin-6-sulfate and lyophilized to form a porous matrix that is further crosslinked for stability. These composites can enhance the growth rate of seeded cells [40]. A recently accepted treatment for severe burns is application of an autologous epidermal layer [cultured epithelial autografts (CEA)] formed by culture of keratinocytes obtained from a small biopsy. Here, the epidermal layer has a greater chance for a “take” when placed on a previously grafted dermal layer than when applied directly to a full thickness wound bed. Autologous keratinocytes have been applied to collagen–GAG dermal substitutes in a variety of ways to create skin replacements that have been tested in porcine full thickness wound models. In one study, 10 days after grafting the collagen–GAG matrix onto the wounds, the newly vascularized dermal substrate was grafted with a CEA that had grown for 3 weeks in vitro. A take rate of 98% was achieved by day 7 [41]. In another study, the keratinocyte culture time was eliminated by seeding freshly disaggregated autologous keratinocytes onto the collagen–GAG matrices before grafting and allowing the cells to develop in a matrix in vivo. A confluent epidermis was formed in 19 days, while the matrix was simultaneously invaded by macrophages, fibroblasts, and endothelial cells [42]. Following this approach has led to the production of Integra™ (Integra Life Sciences,

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Plainsboro, NJ), a commercially available collagen–GAG dermal analog with a silastic film layer. It is used to treat burns, generally in combination with split-thickness autografts or CEAs [43]. Nerve regeneration has also been enhanced with collagen–GAG. Matrix was inserted into tubes spanning 10-mm gaps between the severed ends of sciatic nerves in a rat model [45]. The highly porous collagen–GAG matrix stimulated nerve regeneration to a greater extent than saline-filled controls, as indicated by the number of axons present, their diameters, and their conduction velocities [44]. Further work showed greater numbers of largediameter myelinated axons near nerve termini in collagen–GAG matrix–filled tubes than obtained with autografted nerve implantation [45]. Although collagen type I is the primary collagen in most connective tissues, collagen type II is most prevalent in cartilage. Nehrer et al. [46] recently compared type I and type II collagen–GAG as tissue engineering matrices for the repair of chondral defects. Using an adult canine model, they implanted collagen–GAG sponges seeded with autologous articular chondrocytes into chondral defects created in the trochlea grooves of knees of the animals. Results showed increased formation of reparative tissue with cell-seeded collagen–GAG matrices as compared to matrices alone, and there was a higher percentage of cartilage-like material with the type II matrix as compared to type I [46]. In another recent study, Mueller et al. [47] found that type II collagen–GAG matrices seeded with meniscus cells in vitro enhanced the proliferation of these cells and GAG neosynthesis when compared to the effects elicited by type I collagen–GAG matrices. These data suggest further experimentation to create a possible construct for knee meniscus regeneration [47]. Other modifications to collagen matrices have provided environments to enhance possible bone regeneration. In an effort to provide a matrix that resembles natural bone, composites of nano-hydroxyapatite/collagen [48] or calcium phosphate/collagen [49] have been created very recently. When cultured with rat neonatal calvarial explants, cells migrated into the constructs, proliferated, and deposited mineralizing collagenous matrix [48,49]. D Acellular Dermis Rather than building tissue engineering matrices from individual components de novo, a different approach uses existent scaffolds. For example, human skin may be processed to remove all cellular components but retain ECM components. An acellular dermis may then be seeded with fibroblasts and keratinocytes to create dermal–epidermal composites [50,51]. In addition, when seeded with autologous keratinocytes and grafted onto full thickness wounds, such composites re-epithelialized and led to fibroblast ingrowth and angiogenesis within 2 weeks [52]. In this context, AlloDerm™ (LifeCell, Branchburgh, NJ) is a commercially available split-thickness acellular allograft prepared from human cadaver skin and cryopreserved for off-the-shelf use. It is a nonantigenic complete dermal scaffold that includes elastin, proteoglycans, and basement membrane and has been successful in treatment of burns [53]. E Small Intestinal Submucosa Porcine small intestinal submucosa (SIS) is another natural three-dimensional ECM scaffold that has been successfully used to culture a variety of cell types in vitro [54]. It has also been used for urinary bladder regeneration. In rat and canine models, SIS was applied to

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bladders for reconstruction after cystectomies. At 11 to 15 months, all animals had near normal functional capacities and all three layers of the bladder had regenerated as assessed histologically [55]. F Alginate Other biologically derived materials of nonhuman origin have also been investigated as tissue engineering matrices. Alginate is a calcium-gelling linear polysaccharide obtained from various algae. Chondrocytes suspended in alginate were recently used to repair damaged articular cartilage in a rabbit model [56]. Here, allogenic articular chondrocytes were expanded in culture and suspended in a sodium alginate solution. Full thickness defects created in rabbit femoral condyles were filled with alginate–cell suspension that was then gelled in situ with calcium chloride. This created a construct perfectly inserted into the defect and resulted in regenerated tissue almost indistinguishable from the normal cartilage [56]. Hepatocytes have also been successfully cultured in three-dimensional alginate scaffolds. Porous alginate sponges stimulated seeded cells to aggregate into spheroids and secrete albumin, an indication of normal function [57]. Combining matrix material types, such as proteins and polysaccharides, often creates a substrate that temporarily meets the needs of cells. Thus, modifying alginate gels with fibrin resulted in a biomaterial readily applied to cartilage engineering [58]. When human articular chondrocytes were cultured in alginate, porous fibrin, or alginate/fibrin mixtures, only the mixture promoted formation of cartilage tissue. Separately, the alginate appeared to promote early cell proliferation, while the fibrin component enabled the chondrocytes to maintain their normal phenotype and synthesize cartilage [58]. G Chitin (Chitosan) Chitin, the second most abundant biopolymer on earth, has been evaluated for wound healing, drug delivery, and tissue engineering. Obtained from fungal cell walls and arthropod exoskeletons, chitin is a polymer of N-acetylglucosamine molecules held by 1→4 glycosidic linkages. Chitin is typically extracted from arthropod shells by means of acid/alkali treatment, with approximately 7 tons of crab or shrimp shells producing 1 to 2 tons of chitin [59]. The  linkages allow chitin to form as long straight chains with each glucose residue bonded to the next glucose that has rotated 180°. This orientation promotes fibril formation from parallel chains of chitin. Chitin can be treated with hot sodium hydroxide to hydrolyze acetamido groups from the N-acetylglucosamine to produce chitosan, which dissolves more readily and is easier to use as compared to the native chitin polymer. Chitosan has been found to be biodegradable and nontoxic [60]. It has a molecular weight of 800–1500 kDa [60]. Its properties as a hydrogel have been exploited in designing hemostatic agents [61,62] and drug delivery devices [63]. More recently, chitosan has been suggested as a substrate for tissue engineering, as it is a GAG analog [64]. Since chitosan is polycationic, it has been crosslinked with polyanionic chondroitin sulfate–A (CSA) to form sulfated hydrogels. The CSA–chitosan gels supported the differentiated phenotype of seeded articular chondrocytes as evidenced by round morphology, limited mitosis, and type II collagen and proteoglycan production [64]. Additionally, chitosan–tricalcium phosphate sponges have been prepared as scaffolds for bone formation. When seeded with rat fetal calvarial osteoblastic cells, the cells proliferated in a multilayer fashion and deposited a mineralized matrix [65].

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V DEVELOPMENT OF BIOMIMETIC SYNTHETIC POLYMERS Besides the application of biologically derived material for tissue engineering applications, the last two decades have seen major research efforts directed toward the development of synthetic biodegradable polymers for similar purposes. As noted earlier, the properties that such a polymer must possess are biocompatibility, biodegradability, nontoxicity, and noncarcinogenicity. Each must also have good engineering properties like tensile strength, toughness, solubility in organic solvents, stability in a molten state, and preferably low crystallinity and low glass transition temperature. Aliphatic polyesters of the poly(-hydroxyacids) are perhaps the most common biodegradable synthetic polymers known. The general formula –[–O–CH(R)–CO–]– yields poly(glycolic acid) (PGA) when RH and the chiral polymer poly(lactic acid) when RCH3. The L isomer is naturally metabolized, and therefore the isomer of choice for biomaterials is poly-L-lactic acid (PLLA). Poly(iminocarbonates) provide a relatively new class of biodegradable polymers for tissue engineering applications. A Tyrosine-Based Poly(Iminocarbonates) Poly(iminocarbonates) are hydrolytically labile polymers where the monomers are linked by an iminocarbonate group. The iminocarbonate linkage is structurally derived from the carbonate bond by replacing the carbonyl oxygen with an imino group (Fig. 1). Iminocarbonates were first synthesized by Sandmeyer in 1886 and since then have been investigated by several researchers [66–68]. The early investigations have revealed that simple iminocarbonates are hydrolyzed rapidly in both acidic and basic medium [69,70]. Poly(iminocarbonates) were first synthesized in 1964 by Schminke and coworkers, by reacting di-ols and di-cyanates to form a high molecular weight iminocarbonate polymer [71]. Biodegradability rendered by the pronounced hydrolytic instability of the iminocarbonate linkage was not recognized as a desirable property with potential use in biomedical applications until the work of Langer and Kohn appeared [67,72]. Synthesis of poly(bisphenol A–iminocarbonate) [poly(BPA-iminocarbonate)] and the subsequent characterization of its elastic properties, degradation properties, toxicities, and subcutaneous implantation effects were compared to those of poly(BPA-carbonate), a well-characterized polymer, yielding promising results [72,73]. The mechanical properties of poly(BPA-iminocarbonate) were found to be much higher than commercially available poly(BPA-carbonate) and also higher than polylactates, another class of polymers already accepted for biomedical applications. Poly(BPA-iminocarbonate) could be solvent-cast with many known organic solvents, yielding clear, transparent films. From concentrated solutions of the polymer in

Figure 1 (A) Carbonate and (B) iminocarbonate linkages.

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methylene chloride (30%w/w), fibers could be drawn. Degradation studies revealed the hydrolysis of the polymer in neutral pH occurred by a combination of two different pathways, backbone cleavage in a time period of about 30 days and release of BPA with the formation of the more stable carbonate bond through a period of about 200 days [73]. The use of poly(BPA-iminocarbonate) for drug delivery applications was investigated by loading compression-molded discs of the polymer with p-nitroaniline and Eosin Y as model compounds for low molecular weight hydrophobic and hydrophilic drugs, respectively. The release profiles were found to be dependent on the degree of loading. At high loading (~30%w/w), the release profile was found to be very steep showing quick release, while at low loading (~10%w/w), the release profile was more sustained. Toxicity and tissue compatibility studies were also done using a solvent-cast film of the polymer in a rabbit cornea test. It was concluded from these studies that poly(BPA-iminocarbonate) was a nontoxic and highly tissue-compatible polymer. Through this and additional work, poly(iminocarbonates) were introduced in the research area of various medical applications [68,72,74]. An advance in the field is the synthesis and use of related iminocarbonate polymers from -L-amino acids, naturally obtained monomers [67,75]. Such polymers would be metabolized into naturally occurring amino acids by the body [76], thereby eliciting a favorable biological response. Polymerization of “pseudo”-poly(amino acids) through iminocarbonate bonds, rather than peptide bonds, would vastly improve the processability of the biomaterial without compromising degradation and biocompatibility properties [70,75,76]. The natural amino acid, L-tyrosine, is a major focus of poly(iminocarbonate) investigations for biomaterials. Studies resulted in the discovery of pseudo-poly(amino acids), of which amino acid–based poly(iminocarbonates) became a major area of interest in biomaterial research. Since then, the amino acid–based poly(iminocarbonates) that have become a potential topic of investigation are those based on the natural amino acid, L-tyrosine. Poly(desaminotyrosine–tyrosine–hexyl ester–iminocarbonate) [Poly(DTH-iminocarbonate)] has shown the most promising results toward being used as a potential biomaterial (Fig. 2) [69,70,77]. Besides having the general properties of biocompatibility and biodegradability, poly(DTH-iminocarbonate) has been found to have a low processing temperature (Tg ~ 55°C), allowing the possible incorporation of heat-sensitive drugs into the matrix. Poly(DTH-iminocarbonate) has also been found to be highly ductile and to have a tensile strength comparable to poly(L-lactic acid), the latter being an established biomaterial. The interaction of poly(DTH-iminocarbonate) with soft tissue in a subcutaneous implant site (rat) has indicated a high degree of tissue compatibility [77]. The physicomechanical and chemical characteristics of poly(DTH-iminocarbonate), studied both in vitro [69,70] and in vivo [77] have identified the polymer as a potential biomaterial of the future. In summary, tyrosine-based poly(iminocarbonates) exhibit many excellent characteristics that make them strong candidates for biomaterial applications [66–70,72–75,77]. These include the facts that they are derived from natural amino acids like L-tyrosine and other tyrosine derivatives and that they are biodegradable in both acidic and alkaline conditions into natural metabolites like ammonia, carbon dioxide, and the corresponding amino acid building blocks. In addition, these polymers themselves, as well as their degradation products, have been found to have little or no systemic toxicity, and the polymers have shown considerable lack of immunogenicity. Tyrosine-based poly(iminocarbonates) have high strengths and stiffness and low processing temperatures, which are favorable toward fabricating biomedical devices like orthopedic pins, controlled drug-delivery vehicles, and others. Backbones of certain tyrosine-derived poly(iminocarbonates) can have structures related to certain drugs. An example is the backbone of poly(DTH-iminocar-

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Figure 2 The chemical structures of (A) poly(DTH-iminocarbonate) from (B) desaminotyrosine (DAT) and (C) tyrosine–hexyl ester (Tyr-Hes). bonate), which is related to the drug dopamine. This feature may make drug incorporation in such polymer devices easier. Finally, physicomechanical properties of these polymers can be tuned by altering the pendant ester chain of the tyrosine molecule. B Poly(Ethylene Glycol)–Based Materials Though simple in structure, poly(ethylene glycol) (PEG) has been the focus of much interest in the biomedical world (Fig. 3). It is nontoxic and has been approved by the U.S. Food and Drug Administration for internal consumption [78–80]; PEG is used in large quantities for drug compounding and for a wide variety of cosmetic and personal care products; and PEG-proteins have been cleared for clinical trials in humans. Free PEG administered intravenously to humans is readily excreted through the kidney [81]. Poly(ethylene glycol) is poorly immunogenic, a fact that is crucial to the development of PEG-proteins as drugs [82]. Richter and Akerblom studied antibody formation in humans exposed to PEG and PEG-proteins and found that PEG is poorly immunogenic, while PEG-proteins elicit a mild anti-PEG response [83,84]. Interestingly, this response decreases with increased exposure time and is sufficiently weak to be of no clinical significance.

Figure 3 The chemical structure of poly(ethylene glycol) (PEG).

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As discussed, an initial event after a foreign material is exposed to the blood is protein adsorption onto its surface, followed by a complicated sequence of events. Over the years a substantial amount of research has been performed to improve the biocompatibility of synthetic polymeric materials in contact with blood. However, the precise relationship between the nature of the surface, blood compatibility, and the mechanism of surfaceinduced thrombosis has not yet been completely elucidated. Major factors influencing blood interaction at polymer interfaces are determined by the composition of the polymer surface and the physical and chemical interaction that engage it within the biological environment. One approach taken by many groups is to synthesize nonthrombogenic polymers with surfaces tailored to minimize specific blood interactions such as thrombus formation and platelet recruitment and aggregation as described earlier. Poly(ethylene glycol) has received much attention for applications as a biomaterial interface. The main benefits of PEG are its high solubility in water, favorable chain conformation in water that allows for high mobility, and large excluded volume to repel protein and cell interactions. Furthermore, the simple chemical structure of PEG allows for sophisticated and quantitative end group coupling to enhance the chemical reactivity for surface immobilization [85]. These properties of PEG, especially the dynamic motion and excluded volume, were studied in surface-grafted polymers. Mori et al. [86] immobilized PEG on hydrophobic surfaces and demonstrated that the water content, surface mobility, and volume restriction of the hydrophilic interface are critical factors influencing blood interaction. Poly(ethylene glycol) has also been used in biocompatible materials as a coating or incorporated into a hydrogel [87]. These surfaces are expected to be highly biocompatible because protein adsorption to them is low [87,88]. Both the amount of protein adsorption and the magnitude of other biochemical events, such as platelet adhesion, rapidly decline as the PEG molecular weight rises [86,89]. Poly(ethylene glycol) does not harm active proteins or cells although its amphiphilic nature suggests that it might interact with cell membranes, as evidenced by PEG-induced fusion of cells and liposomes [90–93]. Bittner and colleagues observed that PEG can be used to fuse the severed halves of invertebrate nerve cells [94,95]. Here, morphological continuity was demonstrated by transfer of dyes between the fused segments. This work raises the exciting possibility that PEG might be utilized to fuse severed nerve cells in humans. In a related work, Geron and Meiri have shown that PEG induces fusion of synaptic vesicles and surface membranes of nerve cells [96]. C Polymer Scaffolds Restoration of organ function utilizing tissue engineering technologies often requires the use of a temporary porous scaffold. The function of the scaffold is to direct the growth of cells migrating from surrounding tissue or of cells seeded within the porous structure of the scaffold. The scaffold must therefore provide a suitable substrate for cell attachment, cell proliferation, differentiated function, and, in certain cases, cell migration. The polymeric scaffold can also be used to supply growth factors in cell culture and for improving tissue regeneration in vivo [97]. As noted earlier, polymer scaffolds must possess many key characteristics such as high porosity and surface area, structural strength, and specific threedimensional shapes to be useful as materials for tissue engineering. There are many biocompatible materials that could potentially be used to construct scaffolds, but a biodegradable material is normally desirable since the role of the scaffold is usually only a temporary one. The interaction between the scaffold and the biological milieu must be care-

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fully considered. The adhesion of cells, for example, is often mediated by adsorbed peptides or proteins [98]. Furthermore, the interaction between the scaffold and cells can affect cell function. It has been shown that not only the proliferation, but also other processes such as cell differentiation and cell migration depend on interactions with a polymer surface [98–100]. Biodegradable polymers such as poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA) are used extensively for the controlled delivery of protein and peptide drugs [101], for the manufacture of medical devices [102] and wound dressings [103], and for fabricating scaffolds in tissue engineering [104,105]. A potential shortcoming of the aliphatic polyesters is the adsorption of peptides and proteins frequently observed on the surface of PLA and PLGA [106–108]. In some cases, it may be desirable to suppress nonspecific protein or peptide adsorption to the polymeric device used for cell culture. Once the surface of the polymer is masked, it is possible to functionalize it by binding specific ligands [109–111]. By binding specific adhesion factors (such as fibronectin, for example), the polymer surface will be able to promote the adhesion of specific cells while suppressing the nonspecific protein-mediated adhesion of other cell types. This action is important in the healing of bone defects for instance, where excluding nondesirable cell types from the site of healing is a major challenge [112]. One approach for obtaining biodegradable polymers that permit surface modification is the attachment of a PEG chain to a biodegradable polymer chain such as PLA or PLGA. The hydrophilic chain inhibits protein and peptide adsorption and, consequently, regulates the behavior of cells on the polymer surface. D Hydrogels A number of applications of poly(ethylene glycol) polymers (PEO) have found clinical application and are the basis of commercial products. Vigilon™ (C. R. Bard, Inc., Murray Hill, NJ) is a radiation crosslinked, high molecular weight PEO that swells to a high degree with water and is sold as a sheet for wound covering material (see www.bardmedical.com/ skinwound/products/vigilon.html for product information). Lang and Webster have patented a wound healing laminate in which a hydrophilic open-celled foam is sandwiched between an inner polyurethane open-pore net and an outer continuous layer. The structure prevents the ingress of bacteria and controls the egress of water to between 50–2000 g/m2/24 h for a relative humidity difference of 100 to 10% at 37.5°C [113]. Bactericides such as silver sulfadiazine, chlorhexidine hydrochloride, or povidone iodine could be incorporated into the layer of foamed hydrogel either during or subsequent to its formation [113]. VI SPECIFIC TISSUE ENGINEERING APPLICATIONS A Human Phalanx and Joint Models As discussed, the family of polylactides constitutes a group of polymers that is useful as synthetic, biodegradable materials. Two particular polymers within this group, poly(glycolic acid) and poly-L-lactic acid, have been examined recently with respect to their applicability as scaffold structures for the cell seeding of normal connective tissues [114]. The investigation will be discussed to some degree as it represents a good specific example of tissue engineering. In this context, both PGA and PLLA have been found to be suitable for seeding osteoblasts, chondrocytes, and tenocytes to produce constructs of bone, cartilage, and tendon, respectively [114]. In the studies related to their use in this manner, nonwoven

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Figure 4 A scanning electron photomicrograph of a segment of the mesh of nonwoven polyglycolic acid used for scaffolding of tissue engineered models of human phalanges and joints. Rods of uniform diameter comprise the mesh and form interfiber pores of various sizes. Bar  100 m.

PGA mesh (Albany International, Mansfield, MA) having a diameter of 15 m and pore (interfiber) spaces of approximately 75–100 m was applied (Fig. 4). Its hydrolysis was determined to occur over a shorter time period of a few months compared to that of PLLA, which appears to persist for more than a year. The PLLA held a fiber diameter and pore size similar to those of the PGA mesh. Biodegradation for an intermediate time has been achieved with the development of a crosslinked copolymer of PGA/PLLA, whose persistence has been measured as 35% after 20 weeks of implantation in a nude mouse and 5% after 40 weeks in the same model [114]. In these investigations involving PGA/PLLA scaffolds, fresh bovine periosteum, articular cartilage, and tendon were obtained from normal animals 4–6 weeks old, and the three distinct tissues were formed into constructs having the shape and composition of human phalanges and joints. Preparation of the specific cells has been described [114] and can be briefly detailed as follows. Narrow strips of periosteum removed surgically from radial diaphyses of the animals were wrapped about the central portion of PGA/PLLA scaffolds that were molded into the shape of distal or middle phalanges of the human hand. The cambium layer of each dissected periosteal strip was placed in direct contact with its counterpart copolymer and then sutured to complete a ring of tissue about the scaffold. Articular cartilage chondrocytes were isolated by digesting the tissue with 3% collagenase for 12 h. Cells were cultured in Ham’s F12 medium supplemented with 10% fetal bovine serum (FBS), counted, diluted to 1.5  107 cells/mL, and seeded onto PGA cut into 10  10  2 mm sheets. Tenocytes were isolated by the same digestion procedure used with chondro-

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Figure 5 Scanning electron photomicrograph of polyglycolic acid mesh seeded with bovine chondrocytes for 1 week. At this point in time, the cells have proliferated and secreted a matrix that nearly completely covers the polymer. Bar  100 m.

cytes and these cells were suspended and seeded onto PGA sheets in like manner. The periosteum–PGA/PLLA copolymer constructs were next cultured for a week in M199 medium and chondrocyte–PGA (Fig. 5) and tenocyte–PGA constructs were placed in Ham’s F12 medium for the same time period. All cultures were supplemented with 10% FBS, L-glutamine, penicillin, streptomycin, and ascorbate, and incubations were carried out at 37°C in 5% CO2. Following the week of culture, the three tissue constructs were further engineered surgically to create one of three types of experimental composites, models of either a human distal phalanx, a middle phalanx, or a distal interphalangeal joint, each of increasing complexity [114,115]. Distal phalanx models were formed by suturing a single chondrocyte–PGA sheet to one end of a periosteum–copolymer construct molded to a human distal phalanx. Middle phalanx models were designed by suturing a chondrocyte–PGA sheet to each end of a periosteum–copolymer construct shaped as a human middle phalanx and additionally attaching a tendon–PGA sheet to one of the chondrocyte–periosteum interfaces. Distal interphalangeal joints were designed by suturing chondrocyte–PGA sheets to periosteum–copolymer constructs of the distal and middle phalanx models. The apposing articular surfaces so formed were separated by the insertion of a thin small silicone sheet, and this portion of the construct was wrapped with a tendon–PGA sheet to mimic a joint. These composites were implanted under skin flaps on the dorsal surfaces of nude mice 4–6 weeks old. The animals were sacrificed after 20 and 40 weeks of implantation, and constructs retrieved at these times were examined initially by histological staining [114].

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While the culturing of the various cell types led to the proliferation of osteoblasts, chondrocytes, and tenocytes, there was no development of bone, cartilage, and tendon until the composite constructs were implanted into the nude mice and retrieved some weeks later [114]. By 20 weeks, a period during which biodegradation had reduced the polymer scaffolds by 65% of their original mass, the construct models each demonstrated formation of new tissue and mineralization and other characteristics similar to those of normally developing bone, cartilage, and tendon. The periosteum–copolymer and chondrocyte–PGA sheets comprising distal and middle phalanx constructs of all the models had integrated into well-defined intact osteochondral junctions. The central portions of the periosteum– copolymer constructs were noticeably vascularized both externally and internally. On staining with Safranin-O red, cartilage after 20 weeks of implantation was marked by three distinct zones of tissue, richer in stain intensity with progressively increasing depth from the cartilage surface [114]. Typically, a relatively lightly stained region containing a number of smaller, flattened chondrocytes appeared at the tissue surface; a darker stained region composed of numerous, somewhat rounded cells enclosed in lacunae was observed beneath the surface zone; and a stained region consisting of even more rounded chondrocytes in large lacunae was found in the deepest zone of cartilage. At the same time, the periosteum–copolymer portion of these composites contained trabeculae of new bone, primarily located along the periphery of the constructs. Hypertrophic chondrocytes were found within the copolymer near these trabeculae, a result suggesting that endochondral ossification was leading to bone formation in this tissue. Osteoblasts, many found adjacent to newly secreted seams of osteoid and new bone, were also present within the copolymer. Tenocyte–PGA sheets had formed insertion sites and fibrocartilage where these constructs had earlier been sutured together with cartilage. The developing tendon consisted of numerous tenocytes aligned in a linear fashion and a matrix principally composed of long, parallel arrays of collagen fibers. As noted previously, the morphological features of the three cell types indicated that biological development and cell response throughout the different regions of these engineered composites were similar to those found in normal tissue. After 40 weeks of implantation in mice, the constructs developed further in their appearance and were marked by increasing polymer degradation, mineral formation, vascularization, and additional features [114]. Remarkably, the chondrocytes comprising cartilage at each end of the periosteum–copolymer constructs were found to arrange themselves in rudimentary cell columns resembling an initial organization of chondrocytes in normal epiphyseal growth plates. A more intact junction between cartilage and subchondral bone and between cartilage and tendon appeared in the composites, as did cancellous, cortical, and lamellar bone regions and a putative Haversian system of vascular channels. The development of these types of experimental models of human phalanges and joints is an important application of tissue engineering in the orthopedic arena and a critical advancement toward understanding bone, cartilage, and tendon formation through the use of biodegradable polymers as delivery means for dissociated cells. Though the previous studies [114,115] demonstrate that whole-joint structures can be formed de novo from the dissociated cells in vivo by utilizing polymer scaffolds for maintaining construct shape and composition, problems remain in producing models that are functionally competent. Significant difficulties lie in overcoming immunological rejection of implanted tissues, maintaining the inherent strength of a mineralized biomaterial, and providing response to nerve and muscle impulses and signals. Design of future composites related to those described here will undoubtedly surmount these current obstacles.

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More fundamental, perhaps, to those impediments just noted is a determination of the identity and the fate of the cell types comprising the engineered phalanx and joint composites. In this context, a principal result of the investigations discussed is that the viability of the cells and the composites rests on a vascular supply from the host mice. With time of implantation, it is not clear whether the cells developing and persisting within the constructs are of bovine or murine origin. Recently, additional studies have been initiated in which in situ hybridization has been used to investigate this question [116]. This work employed antisense 35S-labeled oligonucleotide probes specific for bovine aggrecan, type II collagen, osteopontin, biglycan, and bone sialoprotein, examples of molecules whose gene expression and protein synthesis would be expected for bone, cartlilage, and tendon. The probes were hybridized to the bovine model of a middle phalanx, implanted in nude mice, and removed 20 weeks later, as previously described [116]. Results of all hybridization reactions clearly showed that the model was positive for the bovine-specific genes compared to control tissues [116]. A mouse IgG immunoglobulin failed to react when hybridized to the bovine constructs. The mRNA levels for aggrecan (Fig. 6), type II collagen, and biglycan were detected principally over the construct cartilage regions and for osteopontin and bone sialoprotein over both cartilage and bone portions of the model. Thus, osteoblasts and chondrocytes comprising the middle phalanx and present 20 weeks following implantation were of bovine origin. In this circumstance, it can be concluded that a vascular supply provided

Figure 6 Bright field light photomicrograph of a thick (5-m) section of cartilage from a middle phalanx model engineered from bovine tissue and retrieved after 20 weeks of implantation in a nude mouse. The section was hybridized to a bovine-specific aggrecan oligonucleotide probe and shows positive reactivity over nearly all chondrocytes in the area of interest. Bar  20 m.

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by the nude mouse as host to the bovine construct does not alter the initial phenotype of the model [116]. Investigations are continuing of the biological and physical-chemical processes mediating the development, growth, and tissue engineering applications of these and other models of bone, cartilage, and tendon. B Periosteal Explants of Bone Besides the polylactides described, other biodegradable scaffolds have been examined for their own possible orthopedic applications. This includes the use of fibrin glue to provide a natural biomaterial on which to seed cells for their growth and development [117,118]. In this regard, periosteal explants, in part including osteoblasts and their extracellular matrix constituents, from the diaphyses of the radii from newborn calves were minced and mixed with freshly clotting fibrin solutions formed by combining fibrinogen and thrombin [117]. These admixtures of cells/matrix and fibrin were assayed for cell viability over a 7-day period and found to maintain the periosteal cell components, measured in terms of cell migration and total DNA content within the fibrin composite [117]. These data were significantly greater than those determined from the same cells cultured in the absence of fibrin glue [117]. Freshly prepared admixtures were also injected subcutaneously in the dorsal surfaces of 4- to 6-week-old nude mice [117]. Analysis of these implants after up to 12 weeks showed that firm, vascularized nodules developed over this time. The nodules consisted of a mixture of cartilage and bone as early as 3 weeks after implantation, and they became mineralized following their retrieval after 12 weeks. The development of mineral appeared to proceed by means of an endochondral sequence of ossification since cartilage initially developed within the nodules and was then replaced by bone tissue. Osteopontin, a phosphorylated noncollagenous protein associated with aspects of osteogenesis [119,120], was detected in the nodules and progressively increased over the 12-week period. During the same time, the fibrin glue component of the nodules gradually declined through resorptive processes. Control implants in which fibrin glue alone was injected into other mice yielded no nodules, no osteopontin, and no mineral formation at any time point of specimen harvest. Vascularization at the implant site was absent [117]. These studies of fibrin glue suggest that periosteal cell/matrix composites can be successfully transferred from culture conditions to animal models and remain viable utilizing fibrin as a delivery vehicle. In the presence of fibrin, these cells migrate, continue their development along an osteogenic lineage and retain the potential for the induction of new bone at selected transfer sites [117,118]. As discussed previously, fibrin carriers then may be highly effective in providing osteoinductive cell/matrix aggregates to support transplantation, wound healing, reconstruction, and restoration of an orthopedic nature in vivo. C Biodegradable Drug Delivery Devices for Infection Control Although regeneration of lost tissue is of primary importance in any healing situation, infection control is an issue that must never be neglected. Local delivery of antimicrobial agents to the focal site of infection may result in higher therapeutic indices attributable to high local microbial toxicity and low systemic loads [121,122]. The rationale for local therapy is threefold. First, selective toxicity may not always be achieved with conventional antibiotic (AB) treatment. Second, systemic drug use may be unnecessary, unsafe, or contraindicated. Last, the maximum tolerable systemic dose of an antimicrobial agent may not be sufficient because of poor vascularization, the chronic nature of some infections, or, more importantly, microbial resistance. Increasing bacterial resistance to ABs and fewer

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new ABs demand alternative therapies. One such alternative could be the use of existing chemotherapeutic agents that are specifically delivered to the infection site in concentrations at or below their solubility limit, sufficient to control infection but not to result in host toxicity. Choice of the most appropriate AB is dependent upon (1) infection site, (2) degree of AB penetration into the site, (3) type of infecting bacterium, and (4) bacterial sensitivity to the AB [123]. Use of relatively insoluble ABs should increase their penetration by increasing local AB concentration and exposure time. Furthermore, local delivery of poorly soluble ABs results in undetectable systemic load and therefore minimal selective pressure on commensal organisms distal to the infectious focus [124]. The ideal delivery system would be biocompatible, resorbable, easy to use, and inexpensive. It would also release efficacious amounts of drug over a predetermined time frame. Fibrin has several unique characteristics that make it an ideal candidate for delivery of pharmaceuticals and biologics in humans [125]. As already noted, it is readily resorbable and inherently adhesive, recruiting and trapping products in its three-dimensional matrix. Biologic or chemotherapeutic agents can be readily added to the fibrin employed in surgical procedures for their delivery during hemostasis and surgical sealing procedures. Moreover, fibrin permits release of trapped compounds. This action is governed by a diffusion/dissolution mechanism, whereby product is released to the circulation as the fibrin matrix is subjected to fibrinolysis [122]. In this context, Fibrin Sealant (FS) (American Red Cross, Rockville, MD) has been used to deliver (1) demineralized bone and bone morphogenetic proteins to repair bone defects in rats [126], (2) acidic fibroblast growth factor-1 to Teflon shunts for endothelial cell recruitment forming artificial vascular grafts in dogs [127], (3) antiproliferative chemotherapeutic agents in a mouse model of human ovarian cancer [121], and (4) ABs to treat infection [11,122,128,129]. The addition of ABs into this adhesive hemostat initially produced what was thought to be an excellent system for the treatment of wounds [11]. Antibiotic-supplemented FS was subsequently developed to treat experimental osteomyelitis [130], to repair experimental bone defects (e.g., cortical drill holes, heterologous cancellous transplants, and osteotomies [131]), to facilitate wound healing [12], to treat endocarditis [132,133], to repair recto-vaginal and complex fistulas [134], and to treat experimental keratitis [135]. Soluble AB release from FS was found up to 5–7 days with the greatest percentage of material (85%) released in the first 72 h [11,136–139]. Release of ABs over this relatively short time period most likely resulted from the rapid diffusion of the small, ionic molecules [128,139] designed for maximum absorption during oral and parenteral delivery (clinical formulation). However, inadequate delivery kinetics and fear of viral transmission by this human blood product meant that AB delivery from FS was not suitable and it was abandoned in favor of the use of polymethylmethacrylate beads [140], sponge collagen [141], liposomes [142], and plaster of Paris beads [143]. These AB delivery methods were later shown also to be inadequate for delivery of soluble, clinical formulations of ABs [136]. Recently, the use of a less soluble (nonclinical formulation) AB has been documented to retard its release from virally inactivated FS. This has resulted in extended delivery (42 days) in vitro [122], efficacious control of peritonitis without systemic loading of AB [124], and subversion of antibiotic resistance in several bacteria [144]. The following investigation evaluated FS-delivered tetracycline for infection control in the subcutaneous abscess model [145], as modified by Wood [146–148]. Briefly, male Fisher 344 rats (Charles River Breeding Laboratories, Wilmington, MA) were anesthetized with 30 mg sodium pentobarbital per kilogram body weight. The left flank of rats was shaved and disinfected with iodine and alcohol. Abscesses were induced by the subcutaneous injection of Staphylococcus

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Figure 7 Subcutaneous abscess on the flank of a Sprague-Dawley rat 5 days postinfection with Staphylococcus aureus ATCC 27659 (107 CFU) in a 1-mL slurry of 25% sterilized rat cecal contents (50%w/v in Mueller–Hinton broth). aureus ATCC 27659 (107 CFU) in a 1-mL slurry of 25% sterilized rat cecal contents (50%w/v in Mueller-Hinton broth). Tetracycline free-base (300 mg) was suspended in 1mL fibrinogen solution (120 mg/mL, histidine buffer, pH 7.2). One-half milliliter of thrombin (330 IU/mL in 40 mM CaCl2) was combined with 0.5 mL of the tetracycline–fibrinogen suspension through a dual-flow syringe adapter (Duo-flo Dispenser, Hemaedix, Inc., Malibu, CA) and injected subcutaneously several centimeters from the bacterial slurry, 60 min after abscess induction. Parallel control groups received fibrinogen and thrombin components without tetracycline. Animals were observed daily and abscesses measured by calipers to determine the external abscess area. The external area of abscess was defined as the longest diameter multiplied by its perpendicular diameter [145]. Subcutaneous abscesses formed approximately 5 days after induction (Fig. 7) and continued to grow in volume until they spontaneously burst in 3–5 weeks [145]. The effects of tetracycline release from fibrin were evaluated over the initial 2 weeks of abscess development and compared with fibrin (vehicle)-treated controls. Tetracycline released from fibrin over 10 days resulted in a significant decrease in external abscess area as compared with control abscesses (Fig. 8). These results extend previous data (122,124) indicating that antibiotics released from fibrin matrices readily control localized infection. Concern over extended, localized release of ABs prompted additional investigation assessing toxicity. Assessment of FS–tetracycline cytotoxicity was measured by a transepithelial fluorescein permeability assay in vitro [149]. MDCK cells grown in culture develop functional tight junctions and desmosomes that establish a permeability barrier preventing the passage of most molecules except water and some inorganic ions [150]. MDCK cells were cultured in minimal essential medium (Sigma Chemical Co., St. Louis, MO) containing 10% fetal bovine serum and 0.5% lactalbumin (TCM, Sigma) at 37°C in 5% CO2. The stock culture was passed weekly with trypsin-EDTA (Sigma). On the day of assay, 3-mm membrane inserts (Fisher Scientific, Pittsburgh, PA) were placed into wells of 24-well plates (Fisher). Cells were seeded as 1.5  105 in a final volume of 0.5 mL TCM per insert.

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Figure 8 Control of subcutaneous abscess development with long-term subcutaneous release of tetracycline from fibrin. Tetracycline-free base was suspended in a solution of fibrinogen and mixed with thrombin and injected into abscess-bearing rats (). Fibrinogen without antibiotic was mixed with thrombin and injected into abscess-bearing control rats (). When the cultures were confluent, the TCM was removed, the cells rinsed with Hank’s balanced salts solution (HBSS, Sigma), and inserts placed into new wells containing 0.5 mL of the tetracycline-bound fibrin. After 15 min at 24°C without agitation, the insert contents were decanted and cells gently rinsed thrice with 1 mL of HBSS. One-half milliliter of 0.02% sodium fluorescein (Sigma) was added to each insert that was subsequently placed into a new well containing 0.5 mL HBSS. After 30 min at 24°C, the well contents beneath each insert were measured at 490 nm to detect fluorescein. Membranes with cells incubated with distilled water instead of tetracycline, membranes with cells incubated with fibrin alone, and membranes with cells incubated with HBSS served as controls. Triplicate measurements showed that healthy MDCK cells formed desmosomes and tight junctions that normally exclude fluorescein and that tetracycline concentrations released from the fibrin vehicle were not toxic to MDCK cells (Fig. 9) compared with PBS

Figure 9 Effect of tetracycline free-base released from fibrin on transepithelial fluorescein detected at 490 nm. MDCK cells were used to measure cellular cytotoxicity of various tetracycline concentrations released from fibrin.

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(OD490 of 0.003 0.003) and distilled water (OD490 of 1.759 0.122) controls. Of note is the fact that the highest concentration of tetracycline tested exceeded the recommended minimum inhibitory concentration for S. aureus by at least 100-fold. VII CONCLUSIONS Natural wound healing has evolved over the ages to a process that is extremely efficient and effective. However, when injuries are massive, congenital defects are severe, organs are lost or rendered nonfunctional, or co-morbidities create impaired conditions, the natural healing process is inadequate and requires augmentation. Support can now be accomplished in a small way through concerted and collaborative work incorporating a variety of scientific disciplines. Recent studies have demonstrated early achievements in a promising era directed toward the creation of new tissue previously thought irreplaceable. As outlined in this chapter, significant advances have been made in engineering tissues such as skin, bone, cartilage, muscle, tendon, nerve, and cornea, as well as composite systems like a human phalanx and joint. Furthermore, the initial use of natural ECM and protein polymers has been exploited to develop medical devices for advanced hemostasis and drug delivery. Advances in tissue engineering have been made utilizing the functional and informational materials naturally occurring in the native tissue environment. Although cells contain the genetic blueprint to create complex tissues, they cannot function without interactions with an appropriate matrix. The ECM must contain the proper ratios of structural and multiadhesive components (i.e., cellular and molecular adhesion sites) and must allow for cell motility through porosity or low impedance components. It should also possess characteristics that provide for strength, flexibility, biostability, and biocompatibility. The development of tissues in vitro results from utilization of cells, scaffolding polymers, and chemical components that may be derived from cells or from synthetic sources, mimicking natural tissue substrates. Together the cells and scaffold materials can be designed to provide those properties necessary for expansion of the basic combination to a tissue level of hierarchy. Success to date suggests that the approach may someday develop to organ levels. Further experimentation using material combinations possessing a variety of properties conducive to tissue development will eventually allow us to better meet the needs of millions of suffering patients. ACKNOWLEDGMENTS The authors thank Noritaka Isogai, M.D., Ph.D., and Shinichi Asamura, M.D., Ph.D. (Department of Plastic Surgery, Kinki University School of Medicine, Osaka-Sayama, Japan); Ms. Robin Jacquet, Ms. Jennifer Hillyer, and Ms. Jean Zhang (Department of Biochemistry and Molecular Pathology, Northeastern Ohio Universities College of Medicine, Rootstown, OH); and Ms. Rena Mikhail and Susan Chubinskaya, Ph.D. (Department of Biochemistry, Rush-Presbyterian-St. Luke’s Medical Center and Rush University, Chicago, IL) for their respective assistance in creating tissue engineered models of human phalanges and joints and periosteal cell/matrix–fibrin composites, preparing specimens for histological analysis, and analyzing constructs by in situ hybridization. The studies of tissue engineering of human phalanges and joints were supported by grant AR41452 from the National Institutes of Health (to W. J. Landis). The authors also thank Bethany Burkhart (Northeastern Ohio Universities College of Medicine) and Kelly Huebert (Centers for Disease Control and Prevention, Baltimore, MD)

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4 Biocompatibility, Biostability, and Functional Structural Relationships of Biomaterials E. G. Nordström Åbo Akademi University, Turku, Finland

I INTRODUCTION Larry Hench wrote in 1989, “The primary goal of bioceramics research is and must be the development of reliable prostheses for the repair and replacement of diseased and damaged body parts. However, it would be better if it were possible to prevent or delay progression of disease processes, such as arthritis or osteoporosis, that require the use of bioceramics. An important implication of this essay is the fundamental importance of the inorganic origins of genetic and evolutionary pathways. We must understand more fully the biomineralogical factors inherent in disease processes before long-term solutions or prevention will be achieved. Until then, the need for bioceramic prostheses with projected lifetimes of 30–40 years will still be vital in improving the well being of mankind” [1]. The objective of biomaterial research is to produce and to study biomaterials and devices produced thereof, to characterize the host responses, and to assess the epidemiological and health economics issues of product survival. In this context, a biomaterial is a synthetic, natural, or modified natural material intended to be in contact with the biological system. These aspects include the material and biomechanical properties of single- or multicomponent devices, their conditioning, contact duration, and other characteristics, combined with the assessment of their behavior in various biological systems. This also includes retrieval analysis and characterization of the host responses (Ducheyne P., personal communication). Areas that are currently being widely studied by top level biomaterial scientists include biomaterials and tissue engineering; in vitro synthesis of musculoskeletal tissues; bioactive ceramics including hydroxyapatite (HA); bioactive glass and bioactive compos77

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ites; porous metals surface analysis, biocompatibility, electrochemical properties, materials engineering, mechanical properties, and design and stress analysis; orthopedic and dental applications, implant retrieval and analysis and others. The surface treatments for biomaterials have the following goals: to increase biocompatibility and tissue acceptance, to prevent infection, to reduce abrupt interfaces, and to favor biological integration and long-term in-service behavior. Advantages and disadvantages of these methods are analyzed for the main biomaterials categories: metals, synthetic polymers, ceramics, and materials of biological origin. In addition to appropriate physicochemical and mechanical properties, the existence of biologically functional interfaces with living cells is increasingly desired. In the next 20 years, it appears likely that the design of prostheses or implants in actual clinical use will have to be totally or partly reconsidered. Biomaterials and related process engineering are presented in order to obtain optimal surface and structural biocompatibility of implants and devices. Vital–avital composites for tissue engineering, cell culture models, porous ceramics, and degradable polymers are introduced as examples. Emphasis is placed on the conversion of basic research results into clinical applications and on the use of technologies from nonmedical fields in the medical field and vice versa. II BIOCOMPATIBILITY TESTING The establishment of an objective, easy-to-use technique to evaluate tissue irritation in vivo using noninvasive electrical impedance measurements could be one answer. Such a technique would facilitate testing the biocompatibility of various materials and also quantifying skin diseases and other processes involving structural changes. It has been found that irritation of the oral mucosa not clinically or histological discernible could be detected with a simple device based on electrical impedance techniques. Originally, the key problem was the focusing of the probing electrical field in order to minimize artifacts emanating from tissue layers of no interest. The device was then refined and applied to skin testing. It was found that irritation effects far below the limit of the commonly used visual readings could be detected. In this case, it is desirable to exclude from the measurement tissue layers with no diagnostic information, or at least reduce their influence. Thus, the essential steps of the development of a multifrequency depth-selective device are reported. III BIOSTABILITY OF BIOMATERIALS Macrophage is a central cell type in directing host inflammatory and immune processes; hence, its response to biomaterials (i.e., adhesion and giant cell formation) has a direct impact on material biostability and biocompatibility. In a paper by Kao [2], several in vitro and in vivo techniques from previously published results and current investigations are highlighted and presented to demonstrate a means of delineating a part of the complex molecular mechanisms involved in the interaction between biomaterials and macrophages. Complement component C3 was found critical in mediating the initial adhesion of human macrophages on medical-grade polyether urethane ureas. From radioimmunoassay studies, the presence of a diphenolic antioxidant additive in polyether urethane ureas increased the propensity for complement upregulation, but did not affect adherent macrophage density. The subcutaneous cage–implant system was utilized to confirm the role of interleukin-4 in the fusion of adherent macrophages to

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form foreign body giant cells on polyurethanes in vivo. To probe the function/structural relationship of macrophage-active proteins, fibronectin was employed as a model in the formulation of synthetic oligopeptide mimetics. Peptides were grafted onto previously developed, non–cell adhesive polyethylene glycol–based networks. The results indicate that grafted tripeptide RGD sequence supported higher adherent macrophage density than surfaces grafted with other peptides such as PHSRN and PRRARV sequences. However, the formation of foreign body giant cells on peptide-grafted networks was highly dependent on the relative orientation between PHSRN and RGD sequences located in a single peptide. IV THE BONE-BONDING PROCESS The morphology and electrophoretic parameters of the material surface are essential for the bone-bonding process. These phenomena are of chemical and physical origin. In bioactive glass, the gel formed on the surface of glass is a silica gel (Si-gel). In the porous structure calcium ions react with the gel, forming (SiO)–Ca complexes [3]. When the phosphate concentration in the pore is high enough to exceed the solubility product of apatite, the calcium–gel complexes release the calcium, and highly dispersed apatite crystallites are precipitated in the gel. Peltola [4] suggests that in the case of silica monoliths, the results indicate that a great mesopore volume and a wide mesopore size distribution (2–50 nm), preferably over 5 nm, favor the calcium phosphate formation and that a great surface area is not needed. In this case, the smaller the pore size, the greater their number. These results are analogous with the results obtained for the titania coatings and silica fibers. However, these pore structure analyses report on the bulk structure, giving only indirect information about the surface structures and dimensions. In in vitro studies using simulated body fluid (SBF) for immersing bioactive glass, the glass surface reacts and a reaction layer is formed. In hydroxyapatite-based mica composite surface changes like the dissolution of Ca and P, substitution of phosphate by carbonate results in carbonate–hydroxyapatite. However, comparing bioactive glass with hydroxyapatite shows that there are differences between these materials and their bone-bonding mechanisms. Hench and Andersson stated that there are eleven stages in the reaction site on the implant side of the interface with a bioactive glass [5]. The first stage results in leaching and formation of silanols (SiOH) [see Eq. (1)]. There is a rapid exchange of sodium or potassium from the ceramic surface with H and H3O from solution. This stage is usually controlled by diffusion. The second stage is a loss of silica as Si(OH)4 to the solution as a result of breaking of SiMOMSi bonds and formation of SiMOH (silanols) at the glass solution interface [Eq. (2)]. This stage is usually controlled by interfacial reaction. The third stage is the condensation and repolymerization of a SiO2-rich layer on the surface, depleted in alkalis and alkaline earth cations [Eq. (3)]. The fourth stage includes migration of calcium and phosphate groups to the surface through the SiO2-rich layer, followed by growth of the amorphous calcium phosphate–rich film by incorporation of soluble calcium and phosphate from solution. The fifth stage is the crystallization of the amorphous calcium phosphate and incorporation of OH, CO23, or F anions from solution to form a mixed hydroxyl/carbonate/fluorapatite layer. SiMOMNa  H  OH → SiMOH  Na (solution)  OH SiMOMSi  H2O → SiMOH  OHMSi OSiMOH  HOMSiO → OSiMOMSiO  H2O

(1) (2) (3)

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The earlier described reaction kinetics or sequence of reactions are the basis of the bone-bonding property of bioactive glasses. The kinetics of the reaction stages depend on the glass composition. The surface chemical reactions that occur in the surface layer of a bioactive ceramic result in the formation of carbonate–hydroxyapatite. These five reaction stages that occur on the material side of the interface do not depend on the presence of tissues. They occur in distilled water, tris-buffer solutions, or simulated body fluid. LeGeros proposed the five following stages as being responsible for the formation of a strong interface between bone and another bioactive material, dense hydroxyapatite [6]. There are some differences between these materials and their bone-bonding mechanisms. The five stages, according to LeGeros, are (1) acidification of the microenvironment due to the cellular action on the bioactive material; (2) dissolution/precipitation processes resulting in the formation of CO3-apatite (carbonate–hydroxyapatite) intimately associated with an organic matrix similar to bone apatite; (3) production of adhesive proteins and collagen fibrils containing extracellular matrix; (4) simultaneous mineralization of the collagen fibrils and incorporation of the CO3-apatite crystals (originating from the material) in the remodeling new bone; and (5) interdigitation of the mineralized collagen between the host bone and the bioactive ceramic surfaces and within the pores providing the interfacial strength. According to Hench and Andersson [5] the sixth stage in their reaction sequence is the adsorption of biological moieties in the SiO2–carbonate–hydroxyapatite layer, and stage seven is the action of macrophages, followed by attachment of stem cells as stage eight. Stage nine is then the differentiation of these stem cells, and stage ten the generation of matrix. Finally, stage eleven is the mineralization of the matrix. When different bone-bonding mechanisms are to be discussed, a comparison of different bioactive materials and how the physiological processes affect the reaction kinetics is needed. There are accepted hypotheses and it is interesting to discuss the theories in more detail. The use of new surface characterization methods can help to find answers. V BIOCERAMICS Bioceramics is a term introduced for biomaterials that are produced by sintering or melting inorganic raw materials to create an amorphous or a crystalline solid body that can be used as an implant. The final product can be dense or porous. The components of synthetic ceramics are calcium, silica, phosphorus, magnesium, potassium, and sodium and, in some bioactive glasses, boron. Aluminum is avoided because of its toxicity. However, this is not completely true for Al2O3 ceramics and HA/mica composites, where the dissolution of Al, however, could be neglected. Surface bioactive ceramic materials as, e.g., hydroxyapatite ceramics, both dense and porous, show bioactive behavior and tissue bonding [7]. This is considered as true bone bonding when implanted into biological bone. The first person that successfully prepared synthetic dense hydroxyapatite was Aoki. Aoki managed to synthesize hydroxyapatite for medical use in 1972 [8]. However, the first person to discover the bone-bonding phenomena was Hench [9]. Hench made his first discovery of the bioactivity of Bioglass® at about the same time. The chemical composition of Bioglass is in fact not that of glass forming glasses. Glass is considered to form when SiO2  50%. However, the SiO2 of Bioglass is 45%. In 1969 Hench produced the first samples of Bioglass. After this, many types of bioactive ceramics have been developed [10,11]. These materials show surface changes, including dissolution and

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precipitation [12–19] whereas nonbioactive ceramics show neither precipitation nor bone bonding [15,19]. The bone-bonding capacity of different bioactive materials can be divided according to the chemical and crystallographic properties of each material. Therefore, hydroxyapatite that is crystalline reacts in a somewhat different manner than bioactive glass, which is amorphous. Materials such as Bioglass and apatite–wollastonite ceramics can be classified as materials somewhere between apatite and glass. However, both also have a glassy phase present. In other words they are partly or entirely amorphous. The amorphous phase is more readily corroded by body fluids, and the surface of the materials mobilize elements like sodium, potassium, and calcium, which are essential for bioactivity. VI PREPARATION OF MATERIALS A Synthetic Apatite and Calcium Phosphate Minerals Synthetic crystalline calcium phosphate minerals can be of pure apatite type or consist of two or more phases. A pure hydroxyapatite and carbonate–hydroxyapatite can be synthetically prepared if the Ca/P ratio can be controlled. However, synthesis can also result in two phases, i.e., calcium orthophosphate, 3CaOP2O5 (TCP) and HA, 9CaOCa(OH)23P2O5. Nordström et al. have also reported triphasic apatite structure with calcium pyrophosphate, 4CaOP2O5, together with HA and TCP. When these synthetic materials are implanted in a living body they will, after a period of half a year, be seen to be stabilized and have the same Ca/P ratio as that of mature bone [7]. The active exchanges of ions that occur on the surface leads to the changing composition of material and also the delivery of some elements to the new bone that will form at the interface between the material and the osteogenic cells. Nordström et al. have also reported surface substitution of phosphate by carbonate in materials with different calcium phosphate compositions, resulting in carbonate–hydroxyapatite, which is to be found in newly formed bone at the implant–bone interface in the body environment [12–17]. It has also been shown that an apatitic layer can be formed on a ceramic when soaked in a simulated physiological solution that mimics the ion concentration of body fluids (in vitro prefabrication of the apatite layer) [17–19]. B Coral-Derived Apatite Coral-derived apatite (Interpore 200 or 500®, Interpore International, Irvine, CA) is manufactured from natural coral skeleton, in fact a calcium carbonate material, by a hydrothermal exchange reaction in which the calcium carbonate is converted to calcium phosphate. The result is a mixture of hydroxyapatite, Ca5(PO4)3OH, and fluorapatite, Ca5(PO4)3F. This happens because natural coral contains fluoride. In this process the trabecular structure of the coral remains unchanged. The implants derive from the coral species porites. The whole series of apatite that occur comprise Fluorapatite, Ca5(PO4)3F Chlorapatite, Ca5(PO4)3Cl Hydroxylapatite, Ca5(PO4)3OH Carbonate–fluorapatite (francolite), Ca5(PO4,CO3)3F Carbonate–hydroxylapatite (dahllite), Ca5(PO4,CO3)3OH In porous coral-derived apatite materials, the fabrication of the surface is done by using stem cells from bone marrow. In laboratory tests using rat bone marrow it was seen that

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fibrinogens were present before the sixth day. The first osteoblasts were present by the sixth day. The mineralization proceeded until the 12th and 13th day. At that point the whole surface was coated by new mineral. Cultured bone marrow cells from the ilium in flasks and the adherent cells became confluent after 10 to 14 days in all cases. From 3 mL of human marrow aspirate, an average of 7  106 (range: 5–10  106) cells could be obtained. In this primary culture condition, the culture medium (standard medium) did not contain dexamethasone and -glycerophosphate, which were added in the following subculture condition (osteogenic medium); therefore, the culture did not show mineralized areas even after confluence. C Bioactive Glass Among glasses, Bioglass and the glass S53P4 (Abmin Technologies Ltd., Turku, Finland) are the most widely used. However, these glasses are not suitable for glass pebble or glass fiber manufacturing. Therefore, glasses containing increased amounts of MgO and K2O have been developed. The addition of these oxides changes the viscosity flow and the viscosity curve becomes less steep, meaning that the area for the working range becomes wider. Therefore the glass is not sensitive to crystallization. Glass is considered to form when SiO2  50%. However, if we look at the SiO2 content of Bioglass, as already mentioned, it is only 45%. In fact Bioglass is very difficult to heat treat and physically handle as a glass. Andersson et al. have developed bioactive glasses with a SiO2 content typically between 50–60 wt% [20]. Among these glasses the glass S53P4 is the most widely used. However, from this glass it is not possible to manufacture either glass pebbles or glass fibres. Thus, Brink et al. created the glass 13–93, which is a bioactive glass with physical properties that S53P4 lacks [21] Karlsson et al. developed two completely new glass recipes from their experiments with glass S53P4 and glass 13–93. These glasses, 1–98 and 3–98, are very much like the former 13–93 (Table 1). However, they contain more calcium to make them more bioactive. Second, the amount of phosphorus in glass 1–98 was decreased, and boron was added. This, in turn, makes the manufacturing and annealing of the glass much easier. To compensate for the addition of some elements, the amount of magnesium was decreased. After this, a small change in the shape of the glass occurred. However, this change was not crucial for the manufacture of, e.g., small glass spheres. In Table 1, the glass composition of the most common bioactive glasses is presented, including the Table 1 Theoretical Chemical Composition (wt%) of Commonly Used Bioactive Glasses and Synthetic Hydroxyapatite Compared with Natural Bone Glass Bone Hydroxyapatite Bioglass S53P4 13–93 1–98 3–98

SiO2

Na2O

CaO

K2O

MgO

P2O5

B2O3

Developed by

— — 45 53 53 53 55

2 — 25 23 6 6 4

54 57 25 20 20 22 22

— — — — 12 11 9

1 — — — 5 5 5

43 43 5 4 4 2 4

— — — — — 1 1

Natural H. Aoki L. L. Hench O. Andersson et al. M. Brink et al. K. H. Karlsson et al. K. H. Karlsson et al.

Note: F, CO3, and OH are not included in the calculations for natural bone. This affects to some extent the final P2O5 and CaO values 5 wt%, respectively.

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Bioglass made by Hench. Table 1 also includes natural bone composition, as given by Driessens in 1980 and the theoretical composition of a hydroxyapatite with Ca/P  1.67 [22]. VII METAL AND GLASS SURFACES Titanium implants with rough surfaces have been reported to develop hard-tissue contact earlier and to resist removal torque forces better than implants with smooth surfaces [23,24]. Buser et al. compared three different surfaces with unloaded titanium implants. Acid-etched surface demonstrated the best bone apposition (mean bone–implant contact percentage) and the best removal torque values to the implant surface, the next best surface was titanium plasma-sprayed surface. Polished and fine-textured titanium surfaces were clearly less effective [25,26]. Brunette et al. [24] and Rovensky et al. [27] concluded that cells grown on grooved substrata are rounder than cells grown on flat, smooth substrata. A number of cellular properties, including growth [28], secretion of proteins [29], and gene expression [30] are affected by cell shape. All the earlier biological studies of microrough surfaces were carried out on metal surfaces; none of the previous studies were carried out with microrough bioactive glass surfaces. Thus, there was a need to address the effect of surface topography of bioactive glass to biological response. The ingrowth of bone seems to occur in microfissures of the bioactive glass surface [31]. It has been suggested that the bioactive glass surface must first partially dissolve, thereby increasing the concentration of calcium and phosphate ions in the microenvironment and leading to calcium phosphate precipitation. After precipitation, apatite microcrystals form and associate with organic matrix of bone, causing biological growth of bone tissue. This property of the glass is utilized when increasing the macro- and microporosity of the bioactive glass surface. On one hand, according to Grossner-Schreiber et al. [30], surface roughness might enhance the attachment of the bone by growth-regulating proteins. On the other hand, surface area grows with the increasing roughness and this, in turn, increases reaction area and probably increases the number of convenient places for bone formation. Metallic biomaterials are generally used for the replacement of structural components of the human body such as bones, joints, and tooth roots. When they are implanted inside a body, metallic biomaterials may corrode and/or wear, releasing metal ions and debris, which may have toxic effects on tissues and organs. Since it is important for biomaterials to have no toxicity for a living body, a systematic and quantitative evaluation of the cytotoxicity of metallic elements is required for the development of new metallic biomaterials with superior biocompatibility. In a study by Yamamoto et al. in 1998, the cytotoxicity of 43 metal salts was evaluated by the colony formation method using two kinds of cultured cells [32]. The effects of the difference in valence numbers of metallic elements in the salts on cytotoxicity were examined. The cytotoxicity of the salts of metallic elements’ oxo acids was also investigated. As a result, the intensity of metal salts’ cytotoxicity tends to be quite similar between MC3T3E1 and L929 (the correlation coefficient of metal salts’ IC50 is 0.82). The intensity of metal salts’ cytotoxicity depends on the kinds of metallic elements, their chemical states, and concentrations. The IC50 of the highest toxic salt is 1.36  106 mol/L, which differs four orders of magnitude from the IC50 of the lowest toxic salt. K2Cr2O7, CdCl2, VCl3, AgNO3,

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HgCl2, SbCl3, BeSO4, and InCl3 are high toxic salts in which IC50s are smaller than 105 mol/L for both or either of the cell lines. HgCl, Tl(NO3)3, GaCl3, CuCl2, MnCl2, CoCl2, ZnCl2, NiCl2, SnCl2, IrCl4, TINO3, CuCl, RhCl3, Pb(NO3)2, Cr(NO3)3, and Bi(NO3) are relatively high toxic salts in which IC50s are smaller than 104 mol/L for both or either of the cell lines. VIII MATERIAL SURFACE PROPERTIES It has been proved that the surface properties of the implant such as topography affect the biological interactions between bones and implant [33,34]. In an earlier attempt to increase the bioactivity it was shown that it is possible to promote osteoblast function of bioactive glass with surface modification. It has also been shown that preactivation of bioactive glass in tris-buffer facilitates the selective binding of serum proteins, especially fibronectin [35]. Fibronectin is a well-recognized cell adhesion molecule; it contributes to the adhesion of fibroblasts to nonliving surfaces and to neighboring cells [36]. Kornu et al. [37] and ElGhannam et al. [38] demonstrated that adsorption of serum fibronectin to the surface of bioactive glass promotes osteoblast adhesion significantly. Osteoblasts have been shown to be able to migrate across very rough surfaces [39]. When examining the ability of osteoblast-like cells to attach to titanium surfaces, it was found that a higher percentage of cells attached to the rougher surface [40], whereas protrusions of fibroblasts cannot bend over irregularities with an angle of inclination of more than 17° [41]. Evaluation of the attachment of gingival and periodontal ligament fibroblasts and epithelial cells to various titanium surfaces demonstrated that fibroblasts prefer smoother surfaces, and epithelial cells attached only to the smoothest surfaces [33]. Through histological examination, it has been demonstrated that implants roughened by grit blasting exhibit direct bone apposition, whereas smooth implants demonstrate various degrees of fibrous tissue encasement [42]. To further increase the rates of fixation for the bioactive porous and pretreated microrough glass implants, these are soaked in a simulated body fluid. This operation creates amorphous calcium phosphate in the surface layer of the Si-gel and, if the process is continued, it results in small apatite crystals on the surface of Si-gel. Thus, the optimal time for SBF soaking must be studied to find the optimal circumstances for cell attachment and mineralization to begin. IX CHARACTERIZATION METHODS A Microanalysis The study by Valdre et al. in 1995 considers the use of scanning electron microscopy (SEM) and EDS microanalysis applied to the study of numerous mineral-based biomaterials in common use in odontostomatology [43]. The products studied are the following: reabsorbable Dac Blu, nonreabsorbable Dac Blu, nonreabsorbable atomized Dac Blu, nonreabsorbable fine Dac Blu, reabsorbable Biocoral 450, Calcitite 2040-12, Orthogel, Apagen, BTF 65, Calcitite 4060-2, Osprogel, Bio-Oss, Biostite, Ospro-vit, and Merck Hydroxyapatite. By means of SEM it was possible to study the morphology and the microchemistry of the various biomaterials so as to have information concerning their physical and chemical characteristics, such as the crystalline form, the crystalline aggregations, and the space di-

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mensional distribution of the pores, and to check the possibilities of composition variability. All of these factors are fundamental in the evaluation of the functional biocompatibility of a biomaterial once its performance in a biological environment is known. B X-Ray Power Diffraction This study by Valdre et al., also in 1995, considers x-ray power diffraction (XRD) applied to the study of mineral-based biomaterials used in odontostomatology [44]. By means of this method the following materials were analyzed: reabsorbable Dac Blu, nonreabsorbable Dac Blu, nonreabsorbable atomized Dac Blu, nonreabsorbable fine Dac Blu, reabsorbable Biocoral 450, Calcitite 2040-12, Orthogel, Apagen, BTF 65, Calcitite 40602, Osprogel, Bio-oss, Biostite, Osprovit, Merck Hydroxyapatite. These analyses allow the identification of the crystalline phases and the study of the crystallinity and the crystal chemistry of the samples prepared as powder mixtures. This method permits the determination of the physical-chemical and crystalline characteristics of these mineral-based biomaterials formed by powders or transformable into powders. All this information is indispensable for the evaluation of the functional biocompatibility of a biomaterial when its reaction in a biological environment is already known. This method has a great number of advantages against the traditional methods, since marking with solid phases does not destroy the sample, does not modify the physical or chemical characteristics, and affords more information. C Cell–Biomaterial Interaction Cell adhesion and spreading on biomaterials is a key issue in the study of cell–biomaterial interactions. With the development of new disciplines within biomaterials research such as tissue engineering and cellular therapy, information at molecular and structural levels is needed in order to conceive and design biomaterials that elicit specific, functional cell responses. In their study, van Kooten and von Recum in 1999 determined the formation of focal adhesions and fibronectin fibrillar structures by human fibroblasts and human umbilical vein endothelial cells’ adherence to fibronectin-precoated, smooth, and textured silicones as a function of time [45]. Textures consisted of parallel ridges and 0.5-mm deep grooves with a width of 2, 5, and 10 mm. In addition, pillar and well constructs were used. Cells assembled focal adhesions within the first 24 h of adhesion. Fibronectin production and assembly resulted in a dense fibrillar network by day 6. Initial focal adhesion density and size were dictated by the presence of the texture. Topography also influenced initial fibronectin deposition, although the differences did not result in apparent differences in fibronectin networks after 6 days of incubation. Without fibronectin preadsorption, cells did not proliferate on the silicone surfaces. Cells adhered to glass removed all the preabsorbed fibronectin, whereas on silicone they did not. The present study shows that different textures initially give rise to differences in focal contact and fibronectin fibril assembly. The effects of the small, initial in vitro differences on in vivo tissue biocompatibility remain to be studied. The search for a nonthrombogenic material having patency to be used for small diameter vascular graft applications continues to be a field of extensive investigation. The purpose of the present study was to examine whether surface modification of polytetrafluoroethylene (PTFE, Teflon) and polyethylene terephthalate (Dacron) vascular grafts might extend graft biocompatibility without modifying the graft structure. A series of surface coatings were prepared by modifying the argon plasma–treated PTFE and Dacron grafts

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with collagen IV and laminin and subsequently immobilizing bioactive molecules like PGE1, heparin, or phosphatidyl choline via the carbodiimide functionalities [80]. D Fourier Transform Infrared Spectroscopy Surface analysis by Fourier transform infrared (FT-IR) spectroscopy–attenuated total reflectance revealed the presence of new functional groups on the modified graft surfaces. In vitro studies showed that fibrinogen adsorption and platelet adhesion on modified grafts were significantly reduced. This study proposes that surface grafting of matrix components (collagen type IV and laminin) and subsequent immobilization of bioactive molecules (PGE1, heparin, or phosphatidyl choline) changed the surface conditioning of vascular grafts and subsequently improved their biocompatibility. However, more detailed in vivo studies are needed to confirm these observations [80]. X IMPLANT DESIGN Hip joint implants for reconstructive surgery are widely used. One problem is that the implants sometimes fail. This is because the bone cement that is commonly used for fixation has shown itself to be unreliable in the long run, especially among young active people. An alternative to these implants are implants with a surface on which one hopes to induce new bone growth. The bone bonding can be improved by using bioactive glass as a coating of the implant. It has been shown that by using hydroxyapatite as a coating material the surface can be functional. Important advances have been made in the development of biomaterials science and engineering for foot surgery over the past four decades. Implant materials have been separated into two general categories: temporary implants for bone fixation and permanent implants for joint replacement. As presented, however, currently available temporary implants for bone fixation are often left in place permanently, whereas, in the long run, permanent implants for joint replacement cannot realistically be expected to last the lifetime of the average-aged patient and thus are actually only temporary. The benefits and problems of each of these two implant classes were first presented to set the stage for a discussion of possible future directions in the development of new biomaterials that offer the promise of providing improvements for patient care. For bone fixation in foot surgery, the most promising future biomaterials are presented as fully bioabsorbable polymer matrix composites [77]. These implant materials have the potential to provide the initial strength and stiffness of currently used metal alloys without concerns regarding implant removal. With the development of these materials, clinicians and patients will no longer be forced to choose between the risks of implant retrieval and the risks of leaving the implant behind. Current obstacles that must be overcome before these future materials can be introduced for general clinical use are related to improvements in mechanical property durability and degradation product biocompatibility. For joint replacement, tissue engineered viable biomaterials for permanent articular cartilage replacement are presented as the most important of the future biomaterials. If truly permanent joint replacement materials are to be developed, the implants must be able to regenerate and sustain themselves to retain their properties permanently. Living and sustainable tissues are therefore essential if implant properties are to be permanently maintained, because all nonviable materials are subject to eventual irreversible structural breakdown, degradation, and fatigue. Again, many problems remain to be solved before these envisioned future materials can be brought to ac-

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cepted clinical use. However, substantial advances have already been achieved and have demonstrated the feasibility of the development of these materials. Biomaterials science and engineering remains a very challenging and exciting field of research and development. As technology advances, the problems that are faced become more complex and, now more than ever, require interdisciplinary cooperation between molecular and cell biologists, biomaterials scientists and engineers, and clinicians. This is especially true in the relatively new field of tissue engineering. XI BONE SUBSTITUTES When implants fail it is often in connection with a large-scale bone defect or degradation of the bone surrounding the implant. The important role of revised arthoplasty is to heal this bone defect disturbancy. The result of how well the new prosthesis is fastened depends on the healing process of the bone defect. Bone defects are filled by transferred bone mass. Allogenous (from another person) bone is the most widely used for the purpose. The most common allogenous bone transfer is the head of the hip that was removed during the initial operation. This bone is then mashed in a bone mill. For one revised arthoplastic surgery one to four boneheads are needed. Getting sufficient bone of acceptable quality as regards bacteriological and virological cleanliness is a problem. Even if the donors are chosen and tested with care (serological testing and taking one test from the bacteriological sample) there is always some risk that some infections, e.g., hepatitis, HIV, and bacteria infections, can be transferred from the donor to the receiver. The testing of donors and maintainance of a bone bank system is expensive. However, clinically there is no alternative available at the moment. To create a synthetic product that would replace the use of allogenous bone would be decisive progress. In Japan, where the use of allogenous bone is not used for ethical reasons, some preliminary clinical tests have been made to use hydroxyapatite as a filling material in revised arthoplasty of hip prostheses. These results have been promising and show that synthetic bone substitutes can replace the use of allogenous bone in the future. Many types of biomaterials are used as skeletal bone fillers in reconstructive surgery. Attention is paid to hydroxyapatite due to its high biocompatibility with the surrounding tissue. It deals with the testing of new collagen/hydroxyapatite composite material applied to the bone defect on os parietale of rats. The composite material can be prepared from the bovine atelocollagen dispersion and the dispersion of hydroxyapatite. Collagen serves as a matrix in which the particles of hydroxyapatite are anchored. The composite has the advantage that, after saturation with physiological solution, it is compact and can be shaped. The composite material was implanted in the form of plate into six male Wistar rats to the ground bed on the surface of os parietale. The implants were taken out after 4 months. The macroscopic finding of soft tissue, bones, and implants gave evidence of good healing without any adverse reaction. This was also confirmed by the histological observations. Collagen was resorbed and the rest of the material strongly adhered to the bone. The marked osteocytes were present in the zone of the newly formed bone and the dividing line between new and old bone was clear. The experimental results indicate preconditions for the clinical use of this new composite implant material, the structural improvement of which is in progress [78]. A preliminary report has also been made on the use of hydroxyapatite for surgical procedure on a pseudotumor, a rare but serious complication of the bleeding diathesis in patients with inherited bleeding disorders. Surgical or percutaneous treatment and refilling with fibrin sealant was shown to be a successful therapy in a 19-year-old male with severe

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hemophilia B. The pseudotumor, in the upper pad of the left leg, was filled with hydroxyapatite after surgery. The authors suggest that the use of hydroxyapatite is a new and useful option in the surgical treatment of hemophilic pseudotumor [46]. XII DRUG RELEASE In ophthalmology, there is a need for novel degradable biomaterials for e.g., controlled drug release in the vitreous body. These degradable materials should feature both excellent biocompatibility and well-defined kinetics of degradation. In most cases, poly(D,L-lactic acid), or poly(lactic-co-glycolic acid) are used. These materials, however, suffer from some serious drawbacks, since degradation kinetics are difficult to control, especially since socalled burst-degradation occurs. Here follows a description of a set of novel polymeric networks which largely consist of poly(dimethylamino ethyl methacrylate) [poly (DMAEMA)]; these materials are crosslinked via a dimethacrylate molecule that contains two carbonate groups. This system is susceptible to hydrolytic scission. The degradation products do not exert a catalytic effect on the ongoing degradation reaction (i.e., there is no burst effect). The synthesis of three of these materials is described, which differ merely with regard to the crosslinker content. These materials were characterized through dynamic mechanical thermal analysis (DMTA), 1H NMR, and FT-IR spectroscopy, and scanning electron microscopy. The reaction DMAEMA  2-hydroxyethyl methacrylate (HEMA) was studied in detail using 1H NMR spectroscopy, and these experiments revealed that the reaction of DMAEMA and HEMA produces a random (Bernouillian type) copolymer. From this, it was contended that the new materials have a more or less uniform distribution of crosslinks throughout their volume. Structural degradation of the three materials was studied in vitro, at pH 7.4, 9.1, and 12.0. It was found that the materials exhibit smooth hydrolysis, which can be controlled via the crosslink density and the pH, as was expected a priori. It should be noted that degradation of these materials produces nonhydrolyzable, but water-soluble oligo(DMAEMA) and poly(DMAEMA) molecules. Subsequently in vitro studies on the biocompatibility of these materials were performed. The MTT cytotoxicity assay revealed that the materials were cytotoxic to chondrosarcoma cells. This is most probably due to a local increase of the pH due to the basic character of the pending dimethylamino groups. Cytotoxicity remained virtually unchanged after extended washing with water. This indicates that the cytotoxicity is an intrinsic property of the material and was not caused by remnants of free monomer. Cytotoxicity was also seen in cell cultures (human fibroblasts isolated from donor corneas) which were grown in contact with the materials. It is concluded that the new materials have attractive degradation characteristics, but their cytotoxicity makes them unsuitable for applications in ophthalmology [79]. XIII ZETA POTENTIAL AND ELECTROPHORETIC MEASUREMENTS The electrophoretic mobility and/or the zeta potential reflect only the surface conditions of the colloidal particles, which must be completely independent of the colloidal inner composition. Zeta potential negative values are seen in relation to the test colloid carriers of biomolecular monolayer. When a salt concentration as ZnSO4 is increased, a large amount of Zn2 counterions are present in dissolution which are attracted by test colloid and dipolar cationic biomolecules placed on it, giving a place for the formation of the Stern layer, a rigid layer attached on test colloid, and the diffuse layer border, that is to say, the plane of shear. The co-ions SO24 are strongly repulsed by the test colloid due to its large size. The

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counterion concentration decreases exponentially when the distance from such a plane increases, while the co-ion concentration increases in the same way, even reaching equilibrium in the bulk. The extrapolation of this curve must cut the pH value corresponding to a ZnSO4 concentration, which represents the isoelectric point of the analyzed biomolecule. This value is the same value as reported in the literature and is approximately 4.2 pH units. The zeta potential of human enamel is of physiological importance for interactions between enamel surfaces and the surrounding aqueous medium of saliva. The zeta potentials of both enamel and hydroxyapatite have been examined previously by various techniques. In a recent study, Young et al. [47] examined the zeta potential of human enamel and HA using the Coulter DELSA 440, which, using a laser, makes independent Doppler shift measurements of moving particles in an electric field at four different angles, providing advantages over previous techniques. The enamel and HA particles were suspended directly in different phosphate buffers, or first incubated for 2 h in parotid (PS) or whole saliva (HWS) and then suspended in the same buffers. The enamel and HA particles exhibited an overall net surface potential of 15 to 30 mV, depending on the buffer content. Incubation in PS and HWS gave less negative potentials of 8 to 14 mV. In previous studies, the salivary micellelike structures (SMSs) seen in transmission electron spectroscopy (TEM) of parotid saliva were observed to have a zeta potential of 9 mV. The zeta potential determinations in this study support the concept of an adsorption of mostly SMSs to the enamel surfaces, with a change of the zeta potential of the enamel and HA toward that of the SMSs. XIV FUNCTIONAL AND STRUCTURAL RELATIONSHIPS OF BIOMATERIALS Interaction of body fluids containing enzymes of carbonic anhydrase is essential for precipitation of carbonate hydroxyapatite and to accumulate the mineralization or calcification [48]. In addition to small amounts of sodium, magnesium, and potassium, traces of more than 20 other elements are found [49]. Among these zinc has attracted attention because it is believed to both promote and inhibit mineralization. As Vincent noted as early as 1963, zinc does not enter the structure of dahllite or carbonate hydroxyapatite but is present as “part of a metallo-enzyme in the mineralizing mechanism” [50]. One notes that both alkaline phosphatase and carbonic anhydrase are zinc-containing enzymes [51]. Among the trace metals found there can be little doubt that part, if not all, of the sodium, magnesium, and potassium could enter the apatite structure [52]. Nordström et al. described the entering process and concluded that about one site of the calcium occupied sites is substituted by one potassium ion, corresponding to 10 and 5% of the sites by magnesium and 1% of the sites occupied by sodium [18]. Many submicroscopic substances, like proteins, lipoproteins, antibodies, viruses, and cells, can adsorb onto the surface of microscopically visible particles that shape the ceramic materials used as an implant. Carrier particles include such diverse substances as quartz, glass, silica gel, and silica (inert, bioactive, and metals). The mobility of adsorbed particles has considerable practical and theoretical importance. At the practical level, the microscope method, with its concomitant advantages over others, can be used to measure the mobility of submicroscopic substances; pH–mobility curves of proteins are commonly determined in this way. At the theoretical level, it affords some insight into the theory of charge density and zeta potential, , and calculation for large biological particles.

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Carrier particles are coated by allowing them to remain in contact with a solution containing an excess of the adsorbate dissolved in a buffer solution. The extent of adsorption is a function of adsorbate concentration, and a saturation value of the adsorbate concentration is reached beyond which the mobility no longer increases. The mobility of a particle covered with adsorbate is characteristic of the adsorbed material and not of the particle. In a Ph.D. thesis recently presented by Vuola [53], well-documented HA was chosen as a reference material for calcium carbonate (CC) because of their similarity in structure. The limited resorption of HA was confirmed in their study also [54–56], whereas the resorption of coral varied widely. It has been suggested that the main factor in the resorption process is carbonic anhydrase, an enzyme abundant in osteoclasts [57]. The enzyme lowers the pH at the osteoclast–implant interface, dissolving the calcium carbonate matrix [58]. Resorption is most active in the bone–matrix contact areas and proceeds centripetally [59]. Vuola also noted the same resorption pattern with or without the presence of bone tissue [55,56]. Their conclusion was that resorption through carbonic anhydrase activity is not the only, and possibly not even the most important, resorption mechanism. There was no bone formation in control implants [55]. However, the resorption rate was similar to that in the bone marrow group. The control implants at 6 weeks could not be tested due to deformation [56]. Giant cells were found in control implants. However, it could not be determined whether or not these cells were osteoclasts. In the orthotopic site, where abundant osteoclasts are present, the control implants showed resorption at 4 weeks, but all the implants remained, albeit deformed, at 8 weeks. One would expect faster resorption if the carbonic anhydrase enzyme were the main factor in the resorption process. Fricain et al. investigated mouse macrophages and human fibroblasts in vitro and suggested that these cells are capable of phagosytizing coral. Direct contact between these cells and the coral matrix is a prerequisite for the process [60]. In fact, the control implants were already filled with fibrous tissue at 3 weeks, which could explain the similarity in the resorption rate to that in the marrow group. One implant placed to the iliac crest had no tissue ingrowth at 1 year; this implant remained the largest block throughout the follow-up period. They also studied the histology of the one infected block 1.7 years after implantation and found that the coral matrix was completely preserved in areas without tissue ingrowth. In contrast, the block that resorbed the fastest was fully occupied by bone and fibrous tissue at 1 year. In the study by Vuola et al. the 25- g transforming growth factor–beta (TGF-1) group showed less resorption at 8 weeks than did the controls. In general, the structure of the TGF-1–treated implants was much better preserved than that of the coral controls, in which the matrix had partly collapsed. Diminished resorption of CC has not previously been reported in conjunction with TGF- [61]. Transforming growth factor–beta stimulates fibroblast growth but deactivates macrophages [62–64]. The number of macrophages in the TGF-1–treated implants was lower than in the coral controls. On the other hand, empty areas without any fibrotic tissue inside the CC implants could be detected in the 5- and 25- g 3-week groups, implying possible inhibition of fibrotic tissue growth. This apparent discrepancy in diminished fibrotic tissue growth and known fibrotic tissue stimulation by TGF-1 is difficult to explain. It is nevertheless consistent with the diminished resorption of CC. Resorption appears to proceed more rapidly in animals than in humans. When CC blocks were implanted in the cortex of the femur and tibia in pig and sheep, resorption at 1

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month was 64% and 93%, respectively [65]. Complete resorption of CC granules has been observed at 24 weeks in a connective tissue site in pigs [66]. In experiments by Vuola et al., performed in the intramuscular site, the coral implants were deformed at 6 weeks and one implant was completely resorbed in the control group. At 12 weeks the mean total cross-sectional area was 40% of the original [55]. In humans, degradation seems to be slower and 50% of the blocks used to fill the 10mm cranial burr holes resorbed completely at 1 year [67]. In spinal fusions small fragments of natural coral blocks were found after 1 year [68]. When traumatic methaphyseal defects were filled with coral blocks, resorption times of over 4 years have been reported [69]. Calcium carbonate implants placed into the iliac crest were less than 50% of their original size after 2 years. None of the blocks resorbed completely, and one implant was more than 75% of its original size. Vuola concludes in his work that the resorption of CC is unpredictable and the mechanisms are not fully understood [53]. However, the observation concerning HA was that it was not resorbing. The author would like to stress that the unpredictable behavior of CC very much depends on the influence of phosphorus present. If there is phosphorus to satisfy the CC structure completely, the CC tends to become an apatitic structure. Conversely if carbonate is the dominant anion, the final mineral structure is CC. This was shown by Nordström and Karlsson in 1990 [18]. A The Bovine Serum Albumin Isoelectronic Point The coupling of the protein bovine serum albumin (BSA) to these colloidal carrier particles in order to develop diagnostic tests is an important area of research which has been widely studied over the past 40 years. The microelectrophoresis technique (MELFOS; Physics on Disperse Systems Lab. ISPJAE, Cuba) was used for electrophoretic mobility measurements versus both pH and ionic strength. From these data it is possible to determine the isoelectric point and adsorption isotherm of the adsorbed protein (Fig. 1). In many practical situations the value of zeta potential, , obtained from microelectrophoresis experiments is used to characterize the potential at the outer or diffuse part of the electrical double layer and hence is valuable in discussing electrical double layer interaction between the surfaces.

Figure 1 Experimental curves for the BSA biomolecules isoelectric point (pI). (From Ref. 70.)

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Calculation of zeta potential is one of the first steps used to acquire information on the extent to which colloidal particles are charged. B Affinity of Plasma Proteins The affinity of the compounds for erythrocytes reflected their inhibitory potency against the enzyme. Binding to plasma proteins was more dependent on lipophilicity and was stronger for the substituted sulfoamides. Pharmacokinetic studies in rats showed that the unsubstituted sulfoamides with a high affinity for carbonic anhydrase in erythrocytes have longer half-lives and lower clearance values than the substituted sulfonamides, which were more strongly bound to plasma proteins. Bone Gla-proteins (BGP or osteocalcine) and the corresponding cartilage protein chondrocalcine, as well as other noncollagenous proteins, may also act as inhibitors. X-ray crystal structures of carbonic anhydrase II (CAII) complexed with sulfonamide inhibitors illuminate the structural determinants of high affinity binding in the nanomolar regime. The primary binding interaction is the coordination of a primary sulfonamide group to the active site zinc ion. These sites probably act with apatite structure. In the apatite structure the calcium sites are the ones that will incorporate the zinc and in this way a link to a sulfoamino group and carbonic anhydrase is created. According to Klement and Haselbeck, the zinc atoms can occupy half of the calcium sites [71]. In other words five Zn atoms with respect to ten Ca atoms can be exchanged and offer an extremely high ionic bond between the apatite crystal and protein or osteocalcine. The information of surface-bound complex between analyte A, proteins in the bulk, surface-bound ligand B, particle surface sites, can be represented by the scheme ka

A  B ←→ AB kd

(4)

where ka and kd are the rate constants for the formation of complex. Whereas concentration of unoccupied ligand [B] is the difference between the total amounts of ligand on the surface [B0] and the amount of complex [AB], and if the total amount of ligand [B0] is expressed in terms of the maximum analyte binding capacity of the surface, all concentration terms can then be expressed as zeta potential response in millivolts, and the observed rate complex formation, d [AB]   ka [A][B]  kd [AB] dt

(5)

may be written as d   ka C( max  )  kd dt

(6)

where d /dt is the rate of change of zeta potential signal; C is the concentration of analyte; max is the maximum analyte binding capacity in millivolts; and is the zeta potential signal in millivolts at time t. Equilibrium arises when association of analyte with the surface is balanced by dissociation of the surface-bound complex so that the net rate of complex formation is zero: eq   KA eq  KA max (7) C where KA is the equilibrium association constant or affinity constant. A plot of eq/C

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against eq thus gives a straight line from which max, maximum analyte binding capacity, and KA can be calculated. The relation between the change of the zeta potential, due to the adsorption processes, and the concentration of constant temperature can be expressed as / max. By plotting a linear relation, the max (maximum adsorption capacity for BSA biomolecules) and KA (association constant) can be calculated. Effectuating a linear regression the results obtained were KA  4.1  106 M1 and KD  2.1  107 so max  53 mV C Human Marrow Stem Cells It seems that the fundamental phenomenon that leads to tissue bonding is surface transformation, which can influence surrounding cellular activity and result in stimulatory effects on cell differentiation [72]. Ohgushi et al. have previously demonstrated surface-dependent osteoblastic differentiation on bioactive ceramics [73,74]. Even if the surface change is favorable for tissue bonding and promoting cell differentiation, it was not verified until Ohgushi et al. published their results on osteogenetic differentiation of cultured marrow stromal stem cells on the surface of bioactive glass ceramics, whether or not in vitro prefabrication of apatitic layer on the ceramic surface could stimulate cell differentiation [75]. Yoshikawa et al. showed that by using subcultured human marrow cells, active boneforming osteoblasts appeared to an implanted cultured bone/HA construct as early as 1 week later. The HA in this construct was of synthetic coral-derived apatite origin. The synthetic material is produced by treating coral calcium carbonate with phosphorus. The natural coral usually also contains a certain amount of fluoride. Thus, the composition of the end product is very close to that of fluorapatite. After confluence of the primary culture, the cells were subcultured on other dishes in the osteogenic medium, and the scanning electron microscopy of the subcultured on other culture dishes revealed funicular proliferation of spindle cells. Beneath the layer of growing cells, numerous collagen fibers secreted by the cells were observed. The collagen fibers revealed a rough and mineralized surface, and a large amount of noncollagenous mineralized matrix was also observed between the fibers. These findings indicated the osteogenic differentiation of the human marrow cells in the osteogenic medium and showed that the differentiation resulted in in vitro bone formation. For the in vivo bone formation by the human cultured bone/HA constructs, the fresh human marrow cells were cultured in the standard medium and, after confluence, the cells were subcultured on the porous surface of HA in the osteogenic medium. In this subculture condition, as previously described, a thin layer of the in vitro bone was formed on the pore surface. Therefore, the porous framework of HA on which in vitro cultured bone was layered (cultured bone/HA construct) was successfully constructed. To investigate the in vivo osteogenic potential of the cultured bone/HA constructs, the constructs were grafted into nude mice, and all the constructs showed obvious histological bone formation within the pores. In many pores, the HE staining of decalcified specimens revealed thick lamellar bone formation on the pore surface that was considered to originate from the cultured cells. On the surface of the thick layer of bone tissue, many active osteoblasts creating new bone were clearly seen. In spite of the excellent in vivo bone formation in the porous areas of the implanted cultured bone/HA constructs, no cartilage was detected in the porous areas. In the control mice with implants of hydroxyapatite alone (without culture cells), no bone formation was observed. In other words the human marrow–derived cultured/HA construct showed the de novo bone forming capability [76].

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In conclusion, it is clear that many good bioceramics are available and an attempt has been made to explain the fundamental processes of surface interactions. However, there is still much work to be done at the molecular and atomic level before all the parameters affecting bone-bonding mechanisms are fully understood. ACKNOWLEDGMENTS This work was funded by grants from the Technology Development Centre, Finland (TEKES). The work is also a part of the activities of the Åbo Akademi Process Chemistry Group, a National Centre of Excellence appointed by the Academy of Finland. I especially want to thank Dr Heimo Ylänen who through his project, “Further Developing and Investigating the Properties of Bioactive Glass for Clinical Applications,” supported this work financially. REFERENCES 1. Hench L. L. 1989. Bioceramics and the origin of life. J. Biomed. Mater. Res. 23(7):685–703. 2. Kao W. J. 1999. Evaluation of protein-modulated macrophage behavior on biomaterials: designing biomimetic materials for cellular engineering. Biomaterials 20(23–24):2213–2221. 3. Karlsson K. H. 1999. Bone implants—a challenge to materials science. Ann. Chir. Gynaecol. 88(3):226–235. 4. Peltola T. 2000. Nanoscale dimensions and in vitro calcium phosphate formation: studies on sol-gel–derived materials and bioactive glass. Ph.D. thesis, Turku, pp. 93, Ann. Univ. Turkuensis AI 263. 5. Hench L. L., Andersson Ö. H. 1993. Bioactive glass coating. In: An Introduction to Bioceramics, Vol. 1, Hench L. L., Wilson J., Eds, River Edge: Singapore, pp. 41–62. 6. LeGeros R. Z., Daculsi G. 1990. In vitro transformation of biphasic calcium phosphate ceramics: ultrastructural and physico-chemical characterizations. In: Handbook of Bioactive Ceramics, Vol. II: Calcium Phosphate Ceramics. Yamamuro N., Hench L. L., Wilson J., Eds. CRC Press: Boca Raton, FL, pp. 17–28. 7. Nordström E. G., Niemi L., Miettinen J. 1992. Reaction of bone to HA, carbonate-HA, hydroxyapatite  calcium orthophosphate and to hydroxyapatite  calcium ortho- and pyrophosphate. Bio-Med. Mater. Eng. 2(3):115–121. 8. Aoki H. 1994. Medical Applications of Hydroxyapatite. Ishiyaku EuroAmerica: Tokyo. 9. Hench L. L. 1988. Bioactive ceramics. Ann. NY Acad. Sci. 523:54–71. 10. Kokubo T., Ito S., Sakka S., Yamamuro T., et al. 1986. J. Mater. Sci. 21:536. 11. Nordström E. G., Herø H., Jørgensen R. B. 1994. Mechanical properties of hydroxyapatite/mica composite. Bio-Med. Mater. Eng. 4(4):309–315. 12. Legeros R. G., Orly I., Gregoire M., et al. 1991. In: The Bone Biomaterial Interface. Davies J. E., Ed. University of Toronto Press: Toronto, p. 76. 13. Ducheyne P., Radin S., Ishikawa K. 1992. In: Bone-Bonding Biomaterials, Ducheyne P., Kokubo T., van Blitterswijk C. A., eds. Reed Healthcare Communications Publishers: The Netherlands, p. 213S. 14. Radin S., Ducheyne P. 1993. The effect of calcium phosphate ceramic composition and structure on in vitro behavior. II. Precipitation [published erratum appears in J. Biomed. Mater. Res. 1993 27(11):1461]. J. Biomed. Mater. Res. 27:35–45. 15. Radin S., Ducheyne P. 1994. Effect of bioactive ceramic composition and structure on in vitro behavior. III. Porous versus dense ceramics. J. Biomed. Mater. Res. 28:1303–1309. 16. Kokubo T., Ito S., Huang T., et al. 1990. CaP-rich layer formed on high-strength bioactive glass-ceramic A-W. J. Biomed. Mater. Res. 24:331–343.

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5 Biodegradable Hybrid Porous Biomaterials for Tissue Engineering Tetsuya Tateishi and Takashi Ushida The University of Tokyo, Tokyo, Japan Guoping Chen and Toshimi Murata National Institute of Advanced Industrial Science and Technology, Tsukuba, Japan Shuichi Mizuno Brigham and Women’s Hospital and Harvard Medical School, Boston, Massachusetts

I INTRODUCTION As a rapidly emerging technology, tissue engineering holds the potential of a new approach to the repair and reconstruction of tissue and organs damaged by disease and accident [1–3]. There are several factors required for tissue engineering. They include cell source, growth factor, biodegradable porous scaffolds, and mechanical stress. Biodegradable porous scaffolds play an important role in tissue engineering of cartilage as a physical support and also as an adhesive substrate for the isolated chondrocytes. Ideally, the scaffolds used in this application should meet several design criteria. For instance, they should permit cell adhesion, promote cell growth, and allow retention of differentiated cell function, biocompatibility, biodegradability, highly porosity, mechanically strength, and also malleability into desired shapes [4,5]. Generally, three-dimensional biodegradable porous scaffolds can be fabricated from two kinds of biodegradable polymers. One is synthetic and the other is naturally derived polymer. Merits of synthetic polymer such as biodegradable poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymer poly(lactic-co-glycolic acid) (PLGA) include that they are easily processable into desired shapes, possess good mechanical strength, and their degradation periods can also be manipulated. Demerits of PGA- , PLA- , and PLGA99

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derived scaffolds are that they lack cell recognition signals and their hydrophobic property hinders successful cell seeding. On the other hand, naturally derived biodegradable polymers, such as collagen, offers the advantage of specific cell interaction and hydrophilicity, but scaffolds constructed entirely of collagen have poor mechanical strength and are not easy to handle. Hence, hybridization of these two kinds of biodegradable polymer is performed to develop a novel kind of hybrid biodegradable porous scaffold by forming collagen microsponges in the pores of biodegradable synthetic polymer sponge. Chondrocytes in articular cartilage are embedded in a large amount of extracellular matrix, such as proteoglycan and collagen. Proteoglycans are composed of glycosaminoglycan chains covalently linked to a core protein and are highly negatively charged, causing a large amount of water to be sucked into the matrix. This creates a swelling pressure that enables the matrix to withstand compressive force. Articular chondrocytes are exposed to various magnitudes and cycles of pressure by weight bearing. It is conceivable that mechanical factors influence growth, differentiation, and metabolic function of chondrocytes. In this study, we examined the effects of hydrostatic pressure on the synthesis of proteoglycan, one of the major extracellular matrices of cartilage, by direct compression of a liquid phase in the hybrid porous sponge. II BIODEGRADABLE HYBRID POROUS MATERIAL The use of PLGA sponge as a skeleton facilitated formation of the hybrid sponge into desired shapes with high mechanical strength, while the collagen microsponges contributed specific cell interaction [6,7]. A biodegradable hybrid sponge having the advantages of both synthetic and naturally derived polymers was achieved by this hybridization technique, as shown in Fig. 1. Figure 2 shows the fabrication protocol of the hybrid sponge. At first, a porous biodegradable synthetic polymer sponge with an open cellular morphology was prepared by the particulate-leaching technique using sieved sodium chloride or ice particulates as

Figure 1 Biodegradable hybrid biomaterial.

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Figure 2 Preparation procedure of PLGA-collagen hybrid sponge.

porogen materials. Thereafter, the polymer sponge was soaked in a collagen solution, and the collagen solution–containing sponge was frozen at 80°C and freeze-dried. Finally, the sponge was crossliked by glutaraldehyde gas. The PLGA sponge was prepared by adding NaCl particulates to a PLGA solution in chloroform and leaching them out of the dried PLGA/NaCl composite. The PLGA sponge was sectioned with a razor blade and the cross-sections were observed by scanning electron microscopy (SEM) after gold coating, as shown in Fig. 3. Collagen microsponges with interconnected pore structures were formed in the pores of PLGA sponge. The hybrid structure of PLGA-collagen hybrid sponge was further confirmed by detecting elemental nitrogen, which exists in collagen but not in PLGA copolymer, with SEM–electron probe microanalysis (SEM-EPMA). In Fig. 4, the colored spots show the content of elemental nitrogen and the red spots show the highest content. Nitrogen was detected in the microsponges of collagen and the pore surfaces of PLGA sponge pores, but

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b

Figure 3 SEM photomicrographs of a cross-section of (a) PLGA sponge prepared with NaCl particulates ranging in size from 355 to 425 m, and (b) of PLGA-collagen hybrid sponge prepared with 1.0% type I collagen acid solution.

not in the cross-sections of PLGA regions. This result indicates that microsponges of collagen were formed in the pores of PLGA sponge and that the pore surfaces were also coated with collagen. The wettability of a scaffold is very important for successful cell seeding in the threedimensional scaffold. The contact angle of PLGA sponge with water was about 76°. This indicates that the PLGA sponge was relatively hydrophobic. However, the contact angle decreased to about 31° and the wettability increased after hybridization with collagen. The increase of wettability facilitates cell seeding (Fig. 5). It is necessary for the three-dimensional scaffolds to maintain the desired shapes to reserve sufficient free space for the formation of a new tissue when they are used for tissue engineering. The mechanical strength of the hybrid sponge was measured by static tensile and compression mechanical tests. Table 1 shows the results of static compression tests (stiffness) and static tensile tests (ultimate tensile strength and Young’s modulus). The PLGA-collagen hybrid sponge showed higher static compression strength than PLGA and

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a

b

Figure 4 (a) SEM and (b) SEM-EPMA photomicrographs of a cross-section of a PLGA-collagen hybrid sponge prepared as shown in Fig. 1b. collagen sponges both in dry and wet states. In the case of tensile tests, the hybrid sponge also showed higher dynamic and static tensile strength than PLGA and collagen sponges. The hybrid sponge was used for in vitro and in vivo culture of bovine articular chondrocytes. The bovine articular cartilage was minced into 1- to 2-mm pieces and rinsed in cold PBS and sequentially digested by 0.2% collagenese in Ham’s F12 medium supplemented with antibiotics. The scaffold/chondrocytes constructs were cultured in DMEM medium containing 10% fetal bovine serum (FBS) at 37°C and in 5% CO2 atmosphere for in vitro culture study. The scaffold/chondrocytes constructs were first cultured in vitro for

Figure 5 Water contact angle of PLGA and PLGA-collagen films.

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Tateishi et al. Mechanical Properties of PLGA, Collagen, and PLGA-Collagen Hybrid Sponges

Sponge PLGA Dry Wet Collagen Dry Wet PLGA-collagen Dry Wet

Stiffness (N/mm)

Ultimate tensile strength (MPa)

Young’s modulus (MPa)

5.87  0.28 3.88  0.10

0.16  0.00 0.07  0.01

1.09  0.11 0.70  0.02

0.65  0.05* 0.01  0.00**

0.06  0.01* 0.01  0.00**

0.19  0.02* 0.02  0.00**

7.73  0.29* 4.23  0.12**

0.26  0.01* 0.11  0.00**

1.92  0.16* 1.23  0.09**

Note: The average value and the standard deviations are given; n  6. * P  0.01 (versus respective mechanical properties of PLGA sponge at dry state). ** P  0.01 (versus respective mechanical properties of PLGA sponge at wet state).

Figure 6 In vitro three-dimensional culture of articular chondrocytes.

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b

Figure 7 SEM photographs of bovine articular chondrocytes cultured in PLGA-collagen hybrid sponge for (a) 3 days and (b) 6 weeks.

1 week and then implanted subcultaneously in the backs of athymic nude mice, as shown in Fig. 6. Figure 7 shows the SEM photomicrographs of chondrocytes cultured for 6 weeks in vitro. The distribution of cells was spatially uniform throughout the hybrid sponge. The growing chondrocytes and secreted extracellular matrix filled the empty space in the hybrid sponge and decreased the empty space after 6 weeks culture. The photographs of chondrocytes after culture in the hybrid sponge for 4, 6, and 8 weeks are shown in Fig. 8. They appeared glistening white macroscopically. Histological examination of the chondrocytes after 4, 6, and 8 weeks culture using hematoxylin and eosin stains revealed that most of the chondrocytes after 4 weeks culture, and almost all the cells after 6 weeks culture, maintained their phenotypical round morphology. The chondrocytes proliferated, and secreted extracellular matrix increased with the culture time. The results of immunohistochemical staining of type II collagen with anti–type II collagen antibody showed that type II collagen was detected in the extracellular matrix af-

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Figure 8 Hematoxylin and eosin staining of chondrocytes cultured in PLGA-collagen hybrid sponge for various periods.

ter culture for 4 weeks and increased with the culture time. These results suggest the formation of cartilage-like tissue in the hybrid sponge when cultured in vitro. Figure 9 shows the in vitro degradation of PLGA-collagen hybrid sponge. The hybrid sponge gradually dissolved during in vitro cell culture. Its weight decreased to 63.1% of the initial weight after 10 weeks. During a longer culture period, the hybrid sponge will continue to degrade and eventually disappear. Only engineered cartilage containing chondrocytes and extracellular matrix are left behind.

Figure 9 Percentage of weight remaining of hybrid sponge compared to day 0 value as a function of degradation time in DMEM serum medium. Error bars represent means SD for n  3.

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The scaffold/chondrocytes constructs were implanted subcutaneously in the backs of athymic nude mice after 1 week in vitro culture. Each animal received four constructs. Animals were sacrificed and specimens removed after 1, 2, and 4 weeks of implantation. Specimens were evaluated grossly and histologically by means of hematoxylin and eosin stains. Excised specimens revealed that all implants were replaced by cartilage of approximately the same dimensions as the original scaffolds. The implant appeared glistening white macroscopically. Histological examination of the implants using hematoxylin and eosin stains revealed cartilage formation throughout the implants as shown in Fig. 10. It can be concluded that the novel hybrid sponge can serve as an effective biodegradable porous scaffold for cartilage engineering from the results of in vitro and in vivo studies. Use of synthetic polymer

Figure 10 Hematoxylin and eosin staining of chondrocytes cultured in PLGA-collagen hybrid sponge in vivo for 4 weeks. (Top: 50; bottom: 400.)

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sponge as a skeleton facilitated formation of the hybrid sponge into desired shapes with high mechanical strength, while collagen microsponges give the hybrid sponge easy cell seeding and good biocompatibility. III HYDROSTATIC PRESSURE SYSTEM Chondrocytes respond to stress loading, changing in composition, mechanical properties, and structure of articular cartilage. The transduction mechanisms for pressure loading in chondrocytes are unclear, but effects of hydrostatic loading have been examined in vitro with cartilage slices and with monolayers of isolated chondrocytes. Those investigations conclude that cell–matrix interactions influence the effects of cyclic hydrostatic pressure on cellular function [8,9]. Our cyclic hydrostatic pressure loading system is composed of a high pressure culture column containing 3D PLGA-collagen sponges suspended in the medium, a pressure loading pump, and a pressure controller, as shown in Fig. 11. The major differences from previous hydrostatic pressure systems are that our system could continuously perfuse medium and cells were cultured in a 3D hybrid sponge scaffold. After preincubation in a 24-well plate for 3 days, hydrostatic pressure was applied, at no perfusion, and 2.8 MPa with continuous medium perfusion 0.3 mL/min. After cells had been exposed to hydrostatic pressure for 96 h, fixation and immunohistochemical staining were carried out using monoclonal antibiobodies against chondroitin 4-sulfate proteoglycan (C4SPG) and keratan sulfate proteoglycan (KSPG). The contents of chondroitin 4-sulfate proteoglycan and keratan sulfate proteoglycan were quantified by ELISA. Bovine nasal cartilage proteoglycan (BNC-PG) monomer was used as a standard. Figure 12a shows that the accumulation of C4SPG increased by 42% under the

Figure 11 A novel pressure/perfusion culture system. A piston pump, a debubbler, a medium reservoir, and a pressure gauge were installed at the inlet of the column connected to a pen recorder. A back-pressure regulator was installed at the outlet of a column. The collagen sponges were suspended in the column and were applied at levels up to 10 MPa.

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a

b

Figure 12 Accumulation of (a) C4SPG and (b) KSPG under three conditions for 96-h culture: (1) static condition (atmospheric pressure in the 24-well plate); (2) medium perfusion at 0.3 mL/min (atmospheric pressure in the column); (3) continuous hydrostatic pressure at 2.8 MPa and medium perfusion at 0.3 mL/min. Each value represents the mean and SD of six or seven samples (*P  0.01, **P  0.001).

condition of perfusion compared with the static condition. Hydrostatic pressure at 2.8 MPa with perfusion significantly enhanced accumulation of C4SPG (P  0.001). Accumulation of KSPG under conditions of perfusion and pressure/perfusion were increased by 6 and 23%, respectively, compared with static condition, although there was no statistical difference among the production of KSPG under the three conditions, as seen in Fig. 12b. IV CONCLUSIONS From these data, it can be concluded that this 3D PLCA-collagen sponge allows evaluation of the effects of various classes of regulatory chemical and physical factors on chondrocyte that are actively producing matrix of cartilage. A novel kind of hybrid sponge was fabricated by forming microsponges of naturally derived collagen in the pores of a biodegradable synthetic polymer sponge. The hybrid structure was confirmed by SEM and SEM-EPMA. Use of the biodegradable synthetic polymer sponge as a skeleton facilitated formation of the hybrid sponge into the desired shapes with high mechanical strength, while collagen microsponges contributed easy cell seeding and good cell interaction. Cartilage-like tissue formed when chondrocytes were cultured in vitro or in vivo in PLGA-collagen hybrid sponge. Medium perfusion appears to be beneficial for both KSPG and C4SPG synthesis by chondrocytes in cell culture. In addition, application of continuous hydrostatic pressure at 2.8 MPa promoted KSPG and C4SPG accumulation. Further studies are necessary to define optimal magnitude and cycle of hydrostatic pressure for effective enhancement of biosynthesis of extracellular matrices.

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Cartilage has only a small capacity for self-repair. Defects in the joint articular cartilage may progress to osteoarthritis and require total joint replacement. Chondrocytes in high-density culture expressed their typical phenotype so that round or polygonal cells were embedded in a metachromatically staining matrix. While chondrocytes were cultured in low density, they lost proteoglycan and type II collagen expression. The application of hydrostatic pressure with perfusion can be used for in vitro engineered tissue for clinical applications. Enhancement of in vitro matrix synthesis by physical factors such as hydrostatic pressure and medium flow may realize in vitro reconstruction of cartilage and with requisite properties for repair of cartilaginous defects. REFERENCES 1. Langer R., Vacanti J. P. 1993. Tissue engineering. Science, 260:920–926. 2. Nerem R. M., Sambanis A. 1995. Tissue engineering: from biology to biological substitutes. Tissue Eng. 1:3–13. 3. Freed L. E., Grand D. A., Lingbin Z., Emmanual J., Marquis J. C., Langer R. 1994. Joint resurfacing using allograft chondrocytes and synthetic biodegradable polymer scaffolds. J. Biomed. Mater. Res., 28:891–899. 4. Peter S. J., Miller M. J., Yasko A. W., Mikos A. G. Polymer concepts in tissue engineering. J. Biomed. Mater. Res. 43:422–427. 5. Mikos A. G., Sarakinos G., Leite S. M., Vacanti J. P., Langer R. 1993. Laminated three-dimensional biodegradable foams for use in tissue engineering. 14:323–330. 6. Chen G., Ushida T., Tateishi T. 2000. A biodegradable hybrid sponge nested with collagen microsponges. J. Biomed. Mater. Res. 51:273–279. 7. Chen G., Ushida T., Tateishi T. 2000. Hybrid biomaterials for tissue engineering. Adv. Mater. 12:455–457. 8. Murata T., Ushida T., Mizuno S., Tateishi T. 1998. Proteoglycan synthesis by chondrocytes cultured under hydrostatic pressure and perfusion. Mater. Sci. Eng. C6:297–300. 9. Mizuno S., Ushida T., Tateishi T., Glowacki J. 1998. Effects of physical stimulation on chodrogenesis in vitro. Mater. Sci. Eng. C6:301–306.

6 Lactide Copolymers for Scaffolds in Tissue Engineering Shin-Ichiro Morita Gunze Limited, Kyoto, Japan Yoshito Ikada Suzuka University of Medical Science, Mie, Japan

I INTRODUCTION: WHY ARE LACTIDE COPOLYMERS USED? Since the concept of tissue engineering was proposed in the 1980s, it has held great promise for people who have lost tissues or have suffered organ dysfunction, becoming a new option in many medical fields [1]. The objective of tissue engineering is basically to construct biological tissues in vitro or in vivo using autologous, allogeneic, or xenogeneic cells with or without biodegradable scaffolds [2]. In this interdisciplinary field of biomedical engineering, scaffolds, cells, and growth factors are considered to be the three major factors. From a practical point of view, tissue engineering using a scaffold is classified into two categories; cellular and acellular. In the former, cells are seeded in or on a biodegradable scaffold. Cells are attached and proliferated there and secrete an extracellular matrix, finally resulting in the construction of natural tissues. In the latter, a scaffold is implanted into the body without any process of cell seeding. Cells and sometimes capillaries invade the scaffold from around the tissue at the initial step. In both cases, scaffolds gradually degrade and finally disappear, leaving regenerated tissues. One of the most famous biodegradable polymers is polyglycolide (PGA), which has been used for three decades as suture and reinforcement material in surgical operation and whose safety has been proved in many medical applications. However, the application of PGA for other medical devices is extremely limited because of its rapid 111

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degradation. In an attempt to overcome this disadvantage, new biodegradable polymers which possess longer biodegradation periods have been developed. One of them is poly-L-lactide (PLLA), which is now an important biodegradable polymer for fabrication of medical devices that need especially high physical strength. As a substitute for metallic fracture fixation devices, PLLA devices for the same purpose have been widely used, mainly in the area of orthopedics. Since the FDA has approved both PGA and PLLA for clinical use, they have been studied for the object of developing other medical devices. On the other hand, to meet the recent demand of scaffolds for tissue regeneration, many new biodegradable polymers have been proposed in addition to PGA and PLLA. One of the key strategies for the new polymer synthesis is copolymerization. Although many copolymers have been synthesized, the copolymer of lactide and glycolide (PLGA) has been most extensively studied. Moreover, copolymers of lactide and lactone or glycolide and lactone have been investigated by many research groups. Among their copolymers, a copolymer of lactide and -caprolactone [P(LA/CL)] has been investigated for biodegradable polymer scaffolds with an intermediate degradable period between PGA and PLLA for tissue engineering. Chemical structures of PGA, PLLA, PLGA, and P(LA/CL) are shown in Fig. 1. Their physical properties are summarized in Table 1 and the changes of weight and tensile strength remaining in vitro at 37°C are shown in Figs. 2 and 3. More detailed properties and synthesis of PLLA, PGA, and PLGA were reviewed in a recent article [3]. Furthermore, the degradation and mechanical properties of some kinds of lactide copolymers were referred to in articles by Grijpma and Pennings [4,5]. Most of the copolymers have already been confirmed to be safe and have begun to be applied for clinical evaluation. Therefore, it is likely that their potential to be scaffolds for tissue engineering is also very high. In this article, biodegradable polymers, especially, lactide copolymers are reviewed as scaffolds of tissue engineering from a practical point of view.

Figure 1 Chemical structure of synthetic biodegradable polyesters used as scaffolds in tissue engineering.

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Table 1 Physical Properties of Synthetic Biodegradable Polyesters Used as Scaffolds in Tissue Engineering

Poly (glycolide)

Poly (L-lactide)

Tm (C)a Tg (C)b Shape

230 36 Fiber

Tensile strength (MPa) Young’s modulus (GPa) B (%)c

890 (fiber)

PdWO

2–3 months

3–5 years

e Pt50

2–3 weeks

6–12 months

8.4 (fiber)

30 (fiber)

170 56 Fiber, sponge, film 900 (fiber)

8.5 (fiber)

25 (fiber)

Poly (-capro lactone) 60 60 Fiber, sponge, film 10–80 (fiber) 0.3–0.4 (fiber) 20–120 (fiber) More than 5 years —

Copolymer of L-lactide and glycolide (10:90)

Copolymer of L-lactide and -capro lactone (75:25)

Copolymer of of L-lactide and -capro lactone (50:50)

200 40 Fiber

130–150 15–30 Fiber, sponge film

90–120 17 Fiber, sponge, film

850 (fiber)

500 (fiber)

8.6 (fiber)

4.8 (fiber)

12 (film)

0.9 (film)

24 (fiber)

70 (fiber)

600 (fiber)

10 weeks

1 year

6–8 months

8–10 weeks

4–6 weeks

3 weeks

a

Melting point. Glass transition temperature. c Elongation at break. d Period until the polymer mass becomes zero (in saline at 37C). e Period until tensile strength of polymers becomes 50% (in saline at 37C). b

Figure 2 The change of weight remaining in vitro hydrolysis as a function of time (in saline at 37°C).

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Figure 3 The change of tensile strength in vitro hydrolysis as a function of time (in PBS at 37°C). II TISSUES REGENERATED USING LACTIDE COPOLYMERS A Skin Regeneration of skin has the longest history and is the most successful field in tissue engineering. In 1975, Rheinwald and Green suggested a method of culturing epidermal keratinocytes under the presence of 3T3 cells [6] and applied this method for burn patients in 1981 [7]. A further application of this method was shown for patients who had sustained extensive burns on more than 95% of their body surface in 1984 [8]. Since then, many tissue-cultured epiderma were developed and became forerunners of tissue engineered products. In contrast, composite cultured skin equipped with an epidermis and dermis in which some kind of scaffold is necessary has been less actively studied. In terms of a scaffold for the purpose of constructing a dermal layer, a membrane made from collagen and chondroitin-6-sulfate (GAG) was suggested [9]. Boyce and coworkers constructed a composite skin by culturing epidermal keratinocytes on a collagen-GAG dermal substitute [10]. With regard to synthetic polymer scaffolds, Cooper and coworkers used polyglactin-910 (90:10 copolymer of glycolide and lactide) as the dermal substitute to overcome the problem of collagen-GAG substitute [11]. Fibroblasts were seeded on knitted Vicryl (polyglactin) and woven Dexon (PGA) to prepare dermis. Implantation of this cultured skin into mouse induced thin epithelium migration from around the tissue and finally constructed a completed skin. Although these two types of composite cultured skin have also advanced to clinical trials, there has not been any comparative study concerning their predominance as a dermal substitute. B Cartilage Regeneration of cartilage is in a preclinical stage of tissue engineering. The epoch-making clinical trial of chondrocyte transplantation by Brittberg and coworkers showed the possibility of a new remedy for deep cartilage defects [12]. In their procedure, autologous chondrocytes cultured in vitro were injected into the defect of a knee joint, followed by covering with a periosteal flap to prevent leakage of cell suspension. To date, several

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companies in Europe and the United States have started commercial application of this technique. On the other hand, Wakitani and coworkers prepared collagen gel containing allogeneic chondrocytes and transplanted them into the cartilage defect of rabbits [13]. Cartilage regeneration using synthetic polymer scaffolds was reported by Vacanti and coworkers [14]. Fibers of polyglactin-910 covered with coating materials were revealed to be a good template for cartilage regeneration. The advantage of synthetic polymer scaffolds over the injection method of cell suspension or transplantation of chondrocytes in collagen gel is the ability to construct engineered tissues with desired dimensions and shapes [15]. As another suitable scaffold, Honda and coworkers recently reported that a biodegradable sponge made of P(LA/CL) was a good template for cartilage regeneration (Fig. 4) [16]. It was also indicated that this type of scaffold possessed sufficient mechanical strength to guarantee a fixed shape. Meniscal tissue regeneration without cell seeding was also attempted using the same polymer scaffold [17]. It seems important to devise a scaffold with sufficient mechanical strength, especially in the case of constructing earlike cartilage. Further studies are required for regeneration of cartilage of much larger defects. C Blood Vessel Vascular prostheses made of synthetic polymers such as poly(ethylene terephthalate) and expanded polytetrafluoroethylene are widely used in many cardiovascular operations. Since the 1980s, biodegradable vascular grafts have been vigorously studied to overcome the shortcomings of nondegradable vascular grafts. Bowald and coworkers observed that biodegradable materials were good candidates for vascular grafts compared to nondegradable grafts [18]. Dogs died from graft rupture when a double-layer mesh made of polyglactin-910 with length exceeding 15 cm was implanted. However, a rupture did not occur for a 7- to 15-cm graft, and the inner surface was well covered with a smooth endothelial layer. They concluded that a biodegradable meshed tube with a nondegradable outer layer was a good substitute for nondegradable grafts. For the same vascular prostheses, Pham and coworkers studied its compliance changes [19], while Galletti and coworkers coated the biodegradable vascular graft with other polymers to control the degradation

Figure 4 Scanning electron micrograph of a P(LA/CL) (75:25) porous scaffold for cartilage regeneration.

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period [20]. In this attempt, occlusion was sometimes observed, and hence new approaches to vascular graft regeneration using cell seeding have recently been noted in the cardiovascular field. In this case, PGA, PLLA, and some kinds of copolymer were found to be important candidates. Shin’oka and coworkers developed a new scaffold made of a polyglactin woven mesh sealed with a nonwoven PGA mesh for tissue engineered blood vessels, and cells were seeded on this scaffold [21]. The tissue engineered vascular graft maintained pulmonary circulation in a growing lamb model and an increase of diameter over time was observed. Although the results seemed very good, they commented that the stiffness of that material was a problem requiring future improvement. To resolve this problem they developed another scaffold made from an elastic P(LA/CL) porous conduit in which a PGA knitted tube was filled (Fig. 5) [22]. After extensive animal experiments, a clinical trial of this material started in early 2000 [23]. Niklasson and coworkers suggested in their unique studies on blood vessel regeneration in vitro that the use of vascular cells and PLGA or PGA scaffolds was very promising [24,25]. This approach seems to be the tissue engineering of the next generation. As another cardiovascular application, Shin’oka and coworkers used their technique for heart valve regeneration [26,27]. Because of the lack of necessity for nutrient supply, tissue engineering in the cardiovascular field seems to be the most realistic and practical.

A

B

Figure 5 Scanning electron micrograph of a scaffold for tissue engineered pulmonary artery made of a P(LA/CL) (50:50) sponge reinforced with a PGA knitted tube. (A) Surface view; (B) cross-section view.

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D Nerve Nerve guides constructed of synthetic polymers such as silicone showed a high possibility for bridging nerve gaps [28]. Biodegradable guiding filaments such as polyglactin were inserted into the silicone tube to extend the gap [29,30], but the permanently remaining outer layer of the silicone tube caused a chronic inflammatory response. To resolve this problem, biodegradable polymer tubes were designed for nerve regeneration. den Dunnenn and coworkers reported that P(LA/CL) was suitable for this purpose because of its low cytotoxicity and minor foreign body reaction as well as suitable biodegradability [31]. They studied this material by microscopically evaluating for 3 to 10 weeks [32]. A further study revealed that nerve function had been recovered until 52 weeks [33] and 2 years [34]. A similar nerve guide made of this polymer had also been synthesized and characterized by another group [35]. Since a series of the aforementioned study on nerve regeneration seemed to be limited to comparably short gaps, a combination of biodegradable scaffolds and cells was recently attempted. Hadlock and coworkers demonstrated that Schwann cells adhered to a PLGA film [36] and constructed a nerve guide of PLGA with a multiple lumen structure by seeding Schwann cells [37]. They pointed out that this structure was necessary to increase the surface area for Schwann cell migration and adherence. The strategy of growth factor incorporation into a polymer foam seems to be another choice for improving nerve guide regeneration [38]. E Liver Regeneration of metabolic organs such as the liver seems to be more complicated than that of skin and cartilage because sufficient supply of nutrients is inevitably essential for the maintenance of hepatocytes. Vacanti and coworkers used multiple fibers of polyglactin910 as a scaffold of liver regeneration [39]. In their preliminary study, the possibility of regenerating liver was observed, although some inflammatory response by the biodegradable polymer occurred. A porous structure is another choice for scaffolds. Mooney and coworkers fabricated biodegradable sponges of PLLA and PLGA using a particular leaching technique [40]. Their animal experiments revealed that the porosity of a scaffold influenced the rate of fibrovascular tissue ingrowth. Prevascularization is another approach to promote liver regeneration. Mikos and coworkers made sponges of PLLA and PLGA to assess the effect of prevascularization on tissue ingrowth [41]. In hepatocytes transplantation research, Mooney and coworkers infiltrated poly(vinyl alcohol) into PLLA or PLGA sponges to modify the hydrophobicity of the polymers [42]. F Dura Mater Many studies have been attempted on artificial dura maters, since the risk of transmitting the Cruetzfeldt–Jacob disease through human cadaveric dura mater was reported. Nussbaum and coworkers demonstrated that a biodegradable artificial dura mater composed of a polyglactin-910 mesh was a good substitute for cadaver dura mater [43]. A collagenous membrane was formed around the implant 4 weeks postoperatively. After 40 days, the material was almost absorbed and adhesion did not occur between the newly generated tissue and the grafted area. For the same purpose, in 1997 Yamada and coworkers developed a biodegradable dura mater made of a P(LA/CL) film reinforced by a PGA nonwoven fabric (Fig. 6) [44]. Animal experiments revealed that a duralike tissue was reconstructed alongside the material 2 weeks after implantation and the dural substitute disappeared at 24

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Figure 6 Scanning electron micrograph of a biodegradable artificial dura mater. A PGA nonwoven fabric is filled in a P(LA/CL) film. weeks postoperatively. Clinical trials of this material were attempted for 20 patients during neurosurgical operations, and a high degree of efficacy was confirmed [45]. These two approaches for regeneration of dura mater are an acellular type of tissue engineering. G Bone In the field of orthopedics, synthetic biodegradable polymers are widely used as fixation devices such as pins, screws, and plates [46]. Poly(lactide-co--caprolactone) was also investigated for this purpose [47]. With respect to scaffolds of tissue engineering for bone, PLLA, PGA, and inorganic materials such as hydroxyapatite and their composites were well reviewed [48]. As for lactide copolymers, Ishaug and coworkers evaluated some biodegradable films made of PGA, PLLA, and PLGA as the scaffold of osteoblast transplantation [49]. Paste of P(LA/CL) was suggested to be good as a filling material for bone defects [50]. Research on lactide copolymers for bone regeneration is very limited compared to other applications. H Other Tissues As PLGA is the most common biodegradable copolymer applied for tissue engineering, several basic studies using this polymer were also achieved for regeneration of some other tissues. In the field of ophthalmology, Giordano and coworkers assessed the attachment of retinal pigment epithelium cells on several thin films of PLLA and PLGA as temporary supporting materials [51]. Aframian and coworkers cultured a salivary epithelial cell line on biodegradable substrata such as PLLA, PGA, and two types of PLGA to develop an artificial salivary gland [52]. Furthermore, regeneration of the pancreas and intestine was attempted with the same technique for the liver using biodegradable polymer scaffolds [39]. III OTHER LACTIDE COPOLYMERS Improvement of biodegradable polymers is required to construct more suitable scaffolds for tissue engineering. One of the promising strategies is to enhance the cell–polymer in-

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teraction. Some polymer scientists attempted copolymerization of lactide and amino acid. One example is the copolymer of lactic acid and lysine [53–55], which is a copolymer capable of attaching the RGD sequence to the side chain. Copolymerization of lactic acid and aspartic acid was also reported [56]. Another approach for the immobilization of RGD into biodegradable polymers is the incorporation of carboxylic acid residues. For this purpose, Yamaoka and coworkers copolymerized lactic acid and malic acid [57]. Although these surface modifications seem to be useful from the viewpoint of cell attachment, further studies should be performed to characterize other properties that are needed for a scaffold in tissue engineering. Improvement of physicochemical properties of other copolymers, especially a copolymer of ethylene oxide and lactic acid, was studied [58,59]. Rashkov and coworkers reported triblock copolymers with a central poly(oxyethylene) ended at both the sides by PLLA to improve the shortcomings of PLLA [60,61]. Although these copolymers seem to have good properties, further studies are required if they are to be applied for tissue regeneration. IV CONCLUSION The greatest merit of lactide copolymers is the ability to fabricate various kinds of scaffolds, as they have a great variety of degradation periods and physicochemical properties compared to homopolymers of PGA and PLLA. It should be emphasized that the role of lactide copolymers in tissue engineering research will become more important in the future. Following the clinical application of skin and cartilage, clinical trials of tissue engineered blood vessels have recently started using P(LA/CL) copolymer. Regeneration of more complex tissues like small phalanges and whole joints was also reported [62]. Since the 1990s, commercial application of tissue engineered products has started and an enormous market is expected in the near future. Moreover, as many countries have started to allocate enormous government budgets for research on tissue engineering, this new biomedical field will become a standard of medicine in place of organ transplantation in the near future. REFERENCES 1. Langer R., Vacanti J. P. 1993. Tissue engineering. Science 260:920–926. 2. Ikada Y. 2000. Prospectus of tissue engineering. Proceedings of Tenth International Conference on Biomedical Engineering, pp. 13–14. 3. Tsuji H., Ikada Y. 1999. Physical properties of polylactides. Curr. Trends Polym. Sci. 4 (Research Trends, India): 27–46. 4. Grijpma D. W., Pennings A. J. 1994. (Co)polymers of L-lactide. 1: Synthesis, thermal properties and hydrolytic degradation. Macromol. Chem. Phys. 195:1633–1647. 5. Grijpma D. W., Pennings A. J. 1994. (Co)polymers of L-lactide. 2: Mechanical properties. Macromol. Chem. Phys. 195:1649–1663. 6. Rheinwald J. G., Green H. 1975. Serial cultivation of strains of human epidermal keratinocytes: the formation of keratinizing colonies from single cells. Cell 6:331–344. 7. O’Connor N. E., Mulliken J. B., Banks-Schlegel S., Kehinde O., Green H. 1981. Grafting of burns with cultured epithelium prepared from autologous epidermal cells. Lancet 1:75–78. 8. Gallico G. G., O’Connor N. E., Compton C. C., Kehinde O., Green H. 1984. Permanent coverage of large burn wounds with autologous cultured human epithelium. N. Engl. J. Med. 311: 448–451. 9. Yannas I. V., Burk J. F. 1982. Wound tissue can utilize a polymeric template to synthesize a functional extension of skin. Science 215:174–176.

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48. Burg K. J. L., Porter S., Kellam J. F. 2000. Biomaterial developments for bone tissue engineering. Biomaterials 21:2347–2359. 49. Ishaug S. L., Yaszemski M. J., Bizios R., Mikos A. G. 1994. Osteoblast function on synthetic biodegradable polymers. J. Biomed. Mater. Res. 28:1445–1453. 50. Ekholm M., Hietanen J., Lindqvist C., Rautavuori J., Santavirta S., Suuronen R. 1999. Histological study of tissue reactions to -caprolactone–lactide copolymer in paste form. Biomaterials 20:1257–1262. 51. Giordano G. G., Thomson R. C., Ishaug S. L., Mikos A. G., Cumber S., Garcia C. A., LahiriMunir D. 1997. Retinal pigment epithelium cells cultured on synthetic biodegradable polymers. J. Biomed. Mater. Res. 34:87–93. 52. Aframian D. J., Cukierman E., Nikolovski J., Mooney D. J., Yamada K. M., Baum B. J. 2000. The growth and morphological behavior of salivary epithelial cells on matrix protein-coated biodegradable substrata. Tissue Eng. 6:209–216. 53. Barrera D. A., Zylstra E., Lansbury P. T., Langer R. 1993. Synthsis and RGD peptide modification of a new biodegradable copolymer: poly(lactic acid-co-lysine). J. Am. Chem. Soc. 115:11010–11011. 54. Hrkach J. S., Ou J., Langer R. 1994. The development of poly(L-lactic acid-co-L-lysine) for tissue engineering: functionalization and new monomer synthesis. Polym. Prepr. 35:450–451. 55. Barrera D. A., Zylstra E., Lansbury P. T., Langer R. 1995. Copolymerization and degradation of poly(lactic acid-co-lysine). Macromolecules 28:425–432. 56. Hrkach J. S., Ou J., Lotan N., Langer R. 1995. Poly(L-Lactic acid-co-aspartic acid): interactive polymers for tissue engineering. Mat. Res. Soc. Symp. Proc. 394:77–82. 57. Yamaoka T., Hotta Y., Kobayashi K., Kimura Y. 1999. Synthesis and properties of malic acid–containing functional polymers. Int. J. Biol. Macromol. 25:265–271. 58. Chen X., McCarthy S. P., Gross R. A. 1997. Synthesis and characterization of [L]-lactide–ethylene oxide multiblock copolymers. Macromolecules 30:4295–4301. 59. Cohn D., Younes H. 1988. Biodegradable PEO/PLA block copolymers. J. Biomed. Mater. Res. 22:993–1009. 60. Rashkov I., Manolova N., Li S. M., Espartero J. L., Vert M. 1996. Synthesis, characterization, and hydrolytic degradation of PLA/PEO/PLA triblock copolymers with short poly(L-lactic acid) chains. Macromolecules 29:50–56. 61. Li S. M., Rashkov I., Espartero J. L., Manolova N., Vert M. 1996. Synthesis, characterization, and hydrolytic degradation of PLA/PEO/PLA triblock copolymers with long poly(L-lactic acid) blocks. Macromolecules 29:57–62. 62. Isogai N., Landis W., Kim T. H., Gerstenfeld L. C., Upton J., Vacanti J. P. 1999. Formation of phalanges and small joints by tissue engineering. J. Bone Joint Surg. 81-A:306–316.

7 Biodegradable Urethanes for Biomedical Applications Sudha Agarwal, Robert Gassner, Nicholas P. Piesco, Sudhakar R. Ganta University of Pittsburgh, Pittsburgh, Pennsylvania

I INTRODUCTION Engineering of any human tissue requires, as basic components, the scaffold or a delivery matrix, viable cells, and cell signals that are capable of directing the differentiation of cell types in order to create a specific organ or tissue. During regeneration, cells provided by the tissue engineered graft or recruited from the host respond to programed signals in the form of growth factors and genes to recreate the damaged tissue, while the matrix or scaffolding provides the crucial substratum in which the cells grow and differentiate. Thus, the scaffolding is essential to support and augment the initial cell growth and differentiation. To accomplish this, it must have material properties that are conducive for the propagation of cells and production of tissue matrix components. Ideally this scaffold must be made of biocompatible materials to avoid rejection by the host. It is also important that the scaffolding matrix is degradable and provides sufficient initial strength to support the injured tissue, but must degrade over time to allow expansion of regenerating tissue. Furthermore, the degradation products of this polymer must be nontoxic molecules that can be metabolized or excreted readily [1–4]. Finally, the matrix must permit the establishment of an extensive vasculature that can serve as a source of cells, nutrients, and growth factors as well as a vehicle for the removal of waste products [5–8]. Polyurethanes have many attributes that are useful in tissue engineering. Properties such as durability, elasticity, elastomer-like character, fatigue resistance, compliance, and tolerance in the body during healing are the strongest points in favor of polyurethanes. Furthermore, bulk and surface modifications via adjustment of hydrophobic/hydrophilic balance and attachment of biologically active molecules to functional groups within the polyurethane structure are possible. Because of these reasons, polyurethanes are considered 123

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as excellent materials for biomedical devices and are widely used. For the most part, their use has been limited to nondegradable matrices [9–14]. Lately, a great deal of interest has been focused on the synthesis of a new generation of nontoxic biodegradable peptide-based polyurethanes [15–29]. The potential usefulness of these polymers lies in the versatility of polyurethane chemistry, which permits the synthesis of an array of peptide-based polyurethanes from amino acid–based hard segments [such as lysine di-isocyanate (LDI)] and polyols (hydroxyl donors) as soft segments (polyesters and sugars) [24,28,29]. The degradation products of these LDI-based urethane polymers are nontoxic lysine and the hydroxyl donor used. The flexibility of polyurethane chemistry allows the incorporation of various proteins (growth factors, adherence molecules, etc.), which can be released in a biologically active form and in a controlled manner during matrix degradation [30–33]. The diversity of polyurethane structure (cellular, membranous, fibrous, foam), the mechanical properties obtainable, their ability to incorporate bioceramics to vary physical/mechanical properties, as well as their moldability and simplicity of synthesis are important factors in favor of the use of biodegradable polyurethanes in tissue engineering. Finally, their general biocompatibility and tolerance by the host are crucial strengths needed for tissue engineering applications. II CHEMISTRY OF URETHANES Polyurethanes represent a generic class of materials formed by a reaction between a multifunctional alcohol (which can be a monomeric or polymeric polyol), the soft segment, and a multifunctional isocyanate, the hard segment [12,14]. Although the formation of the urethane linkage is the primary polymer-forming reaction, NBC double bonds of isocyanates are highly reactive and can combine with electron donors, electron acceptors, as well as newly formed functional groups, giving versatility to the urethane synthesis (Fig. 1). For example, isocyanates can react with amines to form polyurethane-urea or with carboxylic acid to form amides. Isocyanates have very high affinity for water, and reaction of isocyanates with water leads to production of unstable carbamic acids, which decompose to amines. During this process the carbon dioxide is liberated, resulting in formation of a crosslinked foamed polymer networks (Figs. 1 and 2). Side reactions such as biuret and allophanate formation also create crosslink points that contribute to the ultimate physical properties of the polymer. Allophanate is formed in the presence of excess isocyanate when a urethane group donates an active hydrogen from the nitrogen atom which then reacts with isocyanate (Fig. 2). Whereas biuret formation occurs when a urea group donates an active hydrogen to the isocyanate (Fig. 2). These side reactions also permit binding of proteins during prepolymer synthesis and make polyurethane a versatile molecule for use in tissue engineering. Generally, polyurethanes are synthesized in the presence of a catalyst to promote the reaction between isocyanates and polyols. The most commonly used catalysts are mild and strong bases, such as triamines, sodium hydroxide, and sodium acetate. Tertiary amines like DABCO [1,4-diazo(2,2,2)bicyclooctane] and tin compounds (stannous octoate/caprylate) are also commonly used [12]. By lengthening the polymerization time polyurethanes can be successfully formed in the absence of a catalyst [28,29]. Using the conventional urethane chemistry, the earlier biodegradable polyurethanes were synthesized employing lysine di-isocyanate as a hard segment and polyesters, such as polylactide (PLA), polyglycolide (PGA), or -caprolactone, as soft segments [25,26]. In these polymers use of polyesters provided the biodegradability. We have recently synthe-

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Figure 1 Synthesis of polyurethane. The first step in basic polyurethane synthesis is a reaction between an isocyanate and a hydroxyl group in a diol or polyol. This leads to the formation of a polyurethane prepolymer in the presence of excess isocyanates. In the presence of amine groups isocyanates react with amine groups to form poly(urethane-urea) prepolymers. In the second step, the reaction of isocyanates with water leads to formation of amines with the liberation of carbon dioxide, which produces foam. Free amines then react with isocyanate to form crosslinks of the polyurethane structure.

sized similar polymers but have used LDI as a hard segment and small molecular weight hydroxyl donors like glycerol and glucose as soft segments [28,29]. Briefly, these polymers are synthesized by a three-step procedure. During the first step LDI is prepared from lysine ethyl ester using phosgenation reaction, followed by extensive washing. This procedure gave an LDI yield of greater than 98% [29]. In the second step, an isocyanate prepolymer is synthesized by reacting an excess of LDI with a di- (or multifunctional) hydroxy compound (Fig. 3A). This prepolymer is generally a viscous liquid of low molecular weight, but can be of higher molecular weight depending upon the molecular weight of the hydroxyl donor. The prepolymer is then reacted with an additional hydroxy-functional material or amine-functional material in an aprotic solvent to produce a multiblock copolymer. The isocyanate groups, if in excess concentration, can be further reacted with newly formed

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Figure 2 Reactions of isocyanates with various molecules during polyurethane synthesis. Reaction of isocyanate with (A) hydroxyl groups leads to formation of urethane linkages, (B) amine leads to urea linkages, and (C) carboxyl groups leads to amide linkage formation. (D) Allophanate formation: in the presence of excess isocyanates, the urethane group donates an active hydrogen in the nitrogen which reacts with another isocyanate to form allophanate. (E) Biuret formation: in the presence of excess isocyanates, the urea group donates an active hydrogen to the isocyanate. functional groups or hydroxy-terminated molecules to obtain crosslinked structures (Fig. 3B). The third step in the synthesis of polyurethanes is the reaction of prepolymer with water, generating amines and carbon dioxide. The amines then react with the isocyanate groups to form urea. Due to the carbon dioxide release the resultant product is a foam (Fig. 3A). The crosslinks in these foams are through covalent bonds, which hydrolyze during degradation. The utility of these degradable polymers lies in the production of relatively benign lysine during their degradation in vivo. In most biodegradable urethanes, two bifunctional hard segments, lysine diisocyanate and hexane diisocyanate have been employed (Fig. 4). These diisocyanates can readily incorporate a long list of biocompatible diols and polyols as soft segments. In fact any biocompatible material with two or more hydroxyl groups can be used (Fig. 5). Polyols such as glycerol and glucose have been used to synthesize biodegradable urethanes. The advantage of polymers using glycerol and glucose as polyols is that by changing the stoichiometry of the reactants one can create a polymer with a wide array of physical properties. For example, a polyurethane synthesized with LDI/glucose (5:2) is a tightly

Figure 3 Synthesis of biodegradable urethanes. (A) Schematic representation of the synthesis of LDI-glucose polymer showing synthesis of LDI from lysine ethyl ester. Reaction between LDI-glucose (5:2) leads to the formation of highly crosslinked polymer, whereas LDI-glucose (1:1) leads to formation of soft foam. Cross-section of the matrix network synthesized with LDI-glucose at 1:1 ratio showing large pore sizes in the polymer, and the cross-section of the matrix network synthesized with LDI-glucose at 5:2 ratio showing very small pore sizes as a result of tight crosslinks. (B) Synthesis of LDI-glucose–polyethylene glycol (PEG) polyurethane.

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Figure 4 Structures of two diisocyanates used in the synthesis of biodegradable polyurethanes.

Figure 5 Structures of diols and polyols used in the synthesis of biodegradable polyurethanes.

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crosslinked polymer with a mechanical strength similar to coralline hydroxyapatite, whereas LDI/glucose (1:1) is a pliable foam (Fig. 3A). Furthermore, a prepolymer of LDI/glucose (5:2) can be reacted with a variety of appropriate diols to incorporate rigidity/flexibility in the molecule (Fig. 3B). Diols such as aliphatic polyesters (-caprolactone or lactic/glycolic acid) and polyethylene glycol have been used to form biodegradable urethane systems, with a concomitant variation in physical properties [25,26]. The polymers synthesized with polyesters are hydrophilic and are susceptible to hydrolytic cleavage of the ester linkage and degrade over a period of time. One can generate a copolymer of lysine and lactic acid, which, if suitably end capped, functions as a degradable polyol and can be used in the synthesis of biodegradable urethanes [33]. Random, blocky copolymers of polyethylene glycol (PEG) and aliphatic esters combining the physical properties of each in the final material while maintaining degradability have been generated. They combine the hydrophilicity of polyethylene glycol with the stiffness and degradability of a polyester polyol [34]. Biodegradable polyesterurethanes have also been synthesized as phase-segregated multiblock copolymers, e.g., Degrapol® [21,22]. Their synthesis is based on polyaddition of two microdiols that after phase separation form a crystalline and an amorphous phase with lysine methylester diisocyanate or 2,2,4-trimethyl-hexamethylene diisocyanate (TMDI). The crystallizable macrodiol present in these polymers is ,-dihydroxyoligo[((R)-3-hydroxybutyrate-co-(R)-3-hydroxyvalerate)-block-ethylene glycol]. This segment is then chain extended with PCL-diol (a copolyester of adipic acid and ethylene glycol, diethylene glycol, and butane diol). These polymers were found to be slow degrading but biocompatible in vitro and in vivo. Similar biodegradable elastomeric polyurethanes have been used as nerve guidance channels [36]. Interestingly, the microbial polyester 3-hydroxybutyrate-co-3-hydroxyvalerate (3-PHB) also has been used in slow-release drug delivery, surgical sutures, and bone plates due to the biocompatibility of its degradation products [9]. A number of polyurethanes with other amino acid–based hard segments have been synthesized for temporary implantations [24]. These polymers have been synthesized by diesterification of L-phenylalanine and 1,4-cyclohexane dimethanol to yield a diester diamine. To vary the physicochemical and degradation characteristics, the soft segments used in the synthesis of these polymers were polycaprolactone diol (PCL) or polyethylene oxide (PEO) (Fig. 6). III STRUCTURAL ATTRIBUTES OF URETHANES Due to the options available in selecting the chemistries and molecular weights of various components, polyurethanes can be synthesized with a wide range of physical/mechanical properties. Within the general class of polyurethanes, a diversity of polymer characteristics is possible through (1) variations to the structure of both the hydroxy terminal and isocyanate functional species, (2) alterations in the functionality and the chain length of the two primary species to change the topology of polyurethanes, (3) alterations in mechanical properties from those of an elastomer, to those of a thermoplastic, to those of a thermoset, can be achieved via the use of multifunctional isocyanates. For example, in conventional urethanes, bifunctional toluidine diisocyanate (TDI) and hexamethylene diisocyanate (HDI) are employed to prepare flexible foams and elastomers, while multifunctional isocyanates are used to prepare rigid foams and thermosets. Because polyurethanes are generated via a reaction under mild conditions, these materials can be made as neat formula-

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Figure 6 Chain extenders used in the synthesis of biodegradable polyurethanes. (A) Chain extension with lysine ethyl ester and LDI. (B) Chain extension by phenylalanine and hexane diol showing points of hydrolysis. This chain extender is reacted with hexane diisocyanate to form cyanate linkages with diols or polyols.

tions or as reinforced composite formulations, or these can be preformed and then fabricated [12,14,27]. Finally, polyurethanes are unique in that they are inherently foamable and the porosity during foaming can be controlled. The physical properties of polyurethanes depend upon the extent of crosslinking within the polymer, which can be changed by varying the soft segment (polyols) as well as hard segment in the formulations. Polyols provide matrix flexibility and, therefore, are most frequently used to alter the physical/mechanical strength of the urethanes. Both linear and nonlinear polyols (functionality up to 6) are used. Nonlinear polyols (glucose/fructose) are employed to produce network (thermoset) materials where a greater modulus is required. Linear polyols, on the other hand, are employed in applications where either elastomers or flexible foams are needed. To generate a soft, flexible material, polyethylene glycol (PEG), or a PEG-ester copolymer, of relatively high molecular weight have been used [12]. Degradation of these materials leads to formation of hard segment byproducts plus essentially intact PEG chains, which are relatively nontoxic in vivo [25,26,33]. On the other hand, reduction of the molecular weight and/or increasing the functionality of the soft segment employing glucose or glycerol leads to generation of stiffer polyurethanes [28,29]. Aliphatic polyesters (lac-

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tide/glycolide copolymers or polycaprolactone) or polypeptides have been used to fabricate polyurethane matrices with acceptable physical properties [12]. These polyurethanes are biodegradable due to their susceptibility to hydrolysis at ester linkages. Chain extenders are utilized to modify the physical properties of polyurethanes. Small molecular weight diols/polyols or diamines (phenyl alanine, lysine), employed as chain extenders, increase the size of the hard segment and promote microphase separation (Fig. 6). These chain extenders, by generating crystalline domains which function as crosslink points, increase the stiffness of the polyurethane networks [12,28,29]. The number of carbons in the chain extender also affects the hardness of the polyurethanes. For example, glycol-extended polyurethanes exhibit higher elasticity than diamine based-chain extenders [12]. Recently, Skarja and Woodhouse [24] have synthesized biodegradable polyurethanes using diester chain extenders. In these polymers, a diesterification reaction between L-phenylalanine and 1,4-cyclohexane dimethanol was used to yield a diesterdiamine. The physical characteristics and degradation rates of polyurethanes was varied by polymerization with either polycaprolactone or polyethylene oxide (PEO). IV POROSITY IN POLYURETHANE FOAMS Ideally, biomaterials for tissue engineered scaffolds are those that exhibit foamlike microporous structures. The pores of these scaffolds must be well defined and distributed in an open-pore fashion by a system of interconnecting channels to provide a free flow of nutrients and infiltrating cells. These pores should be large enough to allow attachment and proliferation of diverse cell types responsible for the formation of a functional tissue or organ [13,14,27,36,37]. Polyurethanes can be synthesized to fulfill all of these criteria with regard to porosity of an ideal scaffold. Polyurethanes are inherently foams. The density and porosity of foams can be varied by altering the polymerization conditions. Fig. 7 shows variations in the morphology and porosity of LDI-glucose/glycerol polymers introduced during polymerization. In these biodegradable polyurethanes, the porosity was controlled by altering the polymerization conditions or by changing the ratios of soft and hard segments or by varying the soft segment. For example, complete vacuum-assisted dehydration of LDI-glucose prepolymer results in a transparent membrane, which after hydration showed minimal porosity with pores of 3 to 10 m in diameter (Fig. 7A). On the other hand, polymerization of prepolymers of varying compositions in water resulted in foamlike macroporous scaffolds with different mechanical strengths, shapes, and porosities (Figs. 7B–D). The porosities of all three foams were estimated to be between 72 and 82%. The pores in foams B, C, and D were interconnected to allow continuous flow of nutrients in the scaffold and could be populated with cells of various origins (see following discussion). V COVALENT BINDING OF PROTEINS TO POLYURETHANES The proteins placed on the scaffolding matrix play a critical role in the control of cell proliferation, differentiation, and ultimately their organization in a generating tissue. The salient proteins that contribute to cell behavior and tissue development include cell adherence proteins, growth factors, matrix proteins, and their receptors. In fact, a number of molecules placed at the interface of the implant and host show potential to induce or facilitate organized tissue formation [38–41]. For example, application of a delivery matrix with bone morphogenic proteins (BMP) induces bone regeneration in experimental bone defects

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in animals [42–44] as well as when placed on orthopedic implants [45]; and vascular endothelial cell growth factor (VEGF) augments vascularization in soft tissue. Recently, matricellular proteins have been shown to be important mediators of growth factors, extracellular matrix, and cell signaling pathways [45]. Similarly, delivery of extracellular matrix proteins appears to be crucial for cell attachment and availability of certain growth factors to the cells [45]. These observations have necessitated a need for the synthesis of biomaterials with the ability to covalently bind and deliver biological signaling molecules in vivo. One of the inherent strengths of polyurethane chemistry is its ability to bind proteins covalently. Although attachment of proteins to biodegradable polyurethanes has not yet been reported, a plethora of literature exists showing protein binding to bioresistant urethanes [46–51]. The incorporation of peptides either to the isocyanate functional prepolymer or during the foaming reaction results in a polyurethane with covalently bound peptide to the matrix (Fig. 8). To incorporate peptides an aqueous solution of multifunctional hydroxy-capped material or a multifunctional amine is reacted with prepolymer (assuming appropriate stoichiometry and the desired concentration of peptide on polymer). This results in a foamed polyurethane scaffold, with peptide covalently bound to the matrix (Fig. 8). Covalent incorporation of the protein in this manner has been shown to allow retention of its biological activity (in the case of acetyl choline esterase, 96% retention), while

Figure 7 Gross view of the network matrix synthesized with LDI-glucose and/or glycerol. (A) LDI-glucose polymer synthesized as a transparent membrane by vacuum assisted dehydration of prepolymer and subsequent hydration in water. The pore sizes in this polymer were 3 and 10 m. (B) Foam synthesized with 2.5:1 LDI-glucose and porosity controlled by stirring. (B1 and B2) Cross-section of foam showing homogeneous pore sizes between 0.2 and 0.4 mm. (C) Foam prepared with 1:1:1 molar ratios of glucose/glycerol/LDI, showing pore sizes and voids between 0.02 to 3 mm (C1 and C2). (D) LDI and glycerol polymer with large pore sizes and void areas with sizes between 0.3 to 5 mm (D1 and D2). (E) Femur and a piece of proximal tibia showing the cancellous structure of bone showing interconnected pores (E1) similar to those shown in the LDI-based foams (B and C).

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Figure 8 Schematic representation of covalent binding of proteins to LDI–polyethylene glycol polyurethane. The isocyanate groups of the urethane bind amine group(s) of the protein.

increasing long-term stability (acetylcholine esterase continues to retain activity after 3 years) [30,31,52]. Figure 9 shows incorporation of proteins in LDI–multiarm polyethylene glycol (Shearwater Chem, Huntsville, AL). To visualize the binding of proteins to the polymer, two different prestained proteins—elastin coupled with congo red, MW 69,000 (Sigma, St. Louis, MO), and insulin- chain coupled with a ramazol blue dye, MW 3500—were used (Fig. 9A). The binding of elastin congo red (10 mg/mL) to the polymer was carried out during foaming over a 15-min time interval. The polymer foam was then washed thoroughly with phosphate buffered saline (PBS) and cryosectioned to examine the binding of elastin congo red to the polymer foam (Fig. 9B). In another experiment, the prepolymer was exposed simultaneously on the surface to elastin congo red (20 g/L, 0.9% saline) and to prestained insulin blue (20 g/L, 0.9% saline) in the center for 10 min (scheme shown in Fig. 9D). The cross-section of the polymer in Fig. 9C shows the binding of both elastin congo red and insulin blue to the polymer in a spatially distributed manner, where elastin is bound to the surface of the polymer and insulin to the interior of polymer networks. In these reactions the stoichiometry of the protein incorporation into the polymer is controlled by selective addition of a known number of specific protein molecules per functional group of the crosslinking agent and subsequent mixing and layering to build the polymer blocks (Fig. 9D). The ability of polyurethanes to bind proteins has been exploited extensively for the delivery of proteins in vivo. For example, heparin immobilization on 4,4-methylene dicyclohexyldiisocyanate–based polyurethanes is used to reduce cell adhesion [17,46]. Direct covalent binding or photoimmobilization of laminin nanopeptides sequence (CDPGYGSR–NH2) has been utilized to promote the attachment of target cells such as neuronal cells, endothelial cells, and epithelial cells on the prosthetic devices [35–37,51]. Additionally, films of antibodies attached to polyurethanes have been shown to augment endothelial cell adhesion [53]. Polyurethane grafts have been used to deliver biologically active growth factors, such as fibroblast growth factor–2, to enhance wound healing and vascular endothelial cell growth factor to augment vascularization in vivo [54–56]. Recently, the use of slow-degrading polyester polyurethanes coupled with antigens has been proposed for use in immunizations [32]. While the polyurethanes used in most of the preceding cases were slow or nondegradable polymers, the chemistry used to bind proteins can be applied to biodegradable polyurethanes.

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Figure 9 Covalent binding of proteins to an LDI–polyethylene glycol (PEG) polymer. (A) SDS 10% polyacrylamide gel analysis showing prestained elastin–congo red and insulin–ramezol blue. (B) LDI-PEG polymer coupled elastin–congo red. The prepolymer was reacted with elastin–congo red suspended in phosphate buffered saline (PBS) for 15 min and washed extensively. The dark area represents elastin bound to the polymer. (C) LDI-PEG polymer showing spacial binding of elastin–congo red and insulin–ramezol blue. In this experiment, the polymer surface was coupled with elastin–congo red and core was coupled with insulin–ramezol blue. (D) Schematic presentation of the binding of elastin and insulin shown in (C).

VI BINDING OF LIPIDS AND DRUGS TO POLYURETHANES Immobilization of drugs and their controlled delivery is yet another challenge to be met by biodegradable biomaterials [6]. Synthetic biomaterials that have appropriate strengths, degradation times, permeabilities, and an ability to bind drugs are crucial to regulate cell function in many diseases such as diabetes and hormone deficiencies. In addition to proteins, polyurethanes effectively bind phospholipid head groups through isocyanate linkages. This property has been exploited to limit cell attachment to the polyurethanes. Specifically, binding of phospholipid and phosphorylcholine on polyurethanes has been utilized to inhibit protein adsorption and attachment of platelets, bacteria, and neutrophils on catheters [51,57]. Similarly, glycerophophorylcholine when incorporated as a chain extender in poly(tetramethylene oxide)–based polyurethane has been shown to retard bacterial adhesion in vitro [58].

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Binding of such antibiotics as ciprofloxacin, gentamicin, fosfomycin, and flucoxallinon to polyurethanes [Walopur®, a poly(oxytetramethylene glycol) and diisocyanodiphenylmethane polyurethane] has been successfully achieved by the solvent casting technique [30,59]. These drug-bound polyurethanes exhibit sustained release from the Walopur in a concentration and drug dependent manner, while retaining their antimicrobial activity. These observations suggest that availability of biodegradable and biocompatible polyurethanes is likely to accelerate the applications of these polymers in drug delivery. VII BIODEGRADATION OF POLYURETHANES All urethanes are biodegradable to a certain extent [60–62]. The first polyurethanes used in biomedical applications were hydrolytically unstable because these polymers were synthesized with hydrolysis-prone polyester polyols as soft segments [12,25,43]. The next generation of biomedical polyurethanes were synthesized with polyethers, which were hydrolytically stable, but sensitive to oxidation [63–65]. Since then, much effort has been directed toward synthesis of nondegradable polyurethanes of long durability in vivo, using hydrolysis- and oxidation-resistant polycarbonates as soft segment [12–14,27,43,66]. Nonetheless, with the advent of tissue engineering technology, it is becoming clear that biodegradability is an asset for the application of polyurethanes biomedicine. Polyurethanes degrade by several means. The repetitive chain of the polymer is hydrolyzable, oxidizable, and thermolabile. Additionally, polyurethanes are susceptible to biodegradation by enzymes present in vivo [9,23,64,67,68]. Both soft and hard segment composition plays an important role in the biodegradability of polyurethanes. Incorporation of polyesters during synthesis of polyurethanes increases biodegradability due to their susceptibility to hydrolytic and inflammatory enzymes, such as cholesterol esterase and phospholipase A2 [61,69]. Whereas, the microdomains in the hard segment as well as their structure influence the ability of enzymes to degrade the polyurethanes. Thus, a polymer containing a higher number of hydrolytically stable urea and urethane bonds exhibits a lower degradation rate [65]. Mechanisms of biodegradation assessed by Founder transform infrared (FT-IR) spectroscopy show that urethanes degrade via localized oxidation of the soft segment and hydrolysis of urethane bonds joining hard and soft segments [60]. Several oxidative enzymes (proteinase K, chymotrypsin, and thrombin) break down the polyurethanes in vivo [67,70–72]. Furthermore, cholesterol esterase, an enzyme secreted by activated macrophages, is the most potent cause of the polyurethane biodegradation [71]. In addition to chemical breakage, urethanes have been shown to be degraded by nonenzymatic mediators such as superoxide anions, hydrogen peroxide, and hypochlorous acid released by the inflammatory cells surrounding the polymers [70–72]. The initial biodegradable and biocompatible urethanes utilized polyesters susceptible to hydrolysis for degradability [25,26]. Since then many other soft segments have been utilized to synthesize biodegradable polyurethanes. We have shown that LDI-glucose urethane–urea polymers degrade via hydrolysis of the urethane bonds into lysine, glycerol, ethanol, and CO2 in vitro [28,29]. This LDI-glucose polymer degrades in a linear manner at 37°C and shows a consistent breakdown of polymers at a rate of 1.8 mM of lysine per 10 days (Fig. 10B). It has been shown that LDI-lactide, LDI-lactide-co-glycolide, and LDI-caprolactone polymers degrade twice as rapidly in vivo as in vitro [25,26]. Therefore, we suspect the degradation rate of LDI-glucose polymer to be higher in vivo due to the presence of circulating body fluids and physiologically active cells.

resentation of LDI-glucose matrix degradation into lysine, glucose, ethanol, and CO2. (B) Liberation of lysine during degradation of LDI-glucose polymer in PBS over a period of 60 days. LDI-glucose polymer (10 mg/mL PBS) was incubated at 37°C, and 1-mL sample was retrieved every 24 h to examine the release of lysine by Ninhydrin reagent.

Figure 10 Degradation of LDI-glucose polyurethane in aqueous solution. Degradation of LDI polymer in aqueous solution. (A) Schematic rep-

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Other amino acid–based hard segments for temporary implantation include polymers synthesized with an aliphatic diisocyanate, diester chain extenders, and polyesters as soft segments (e.g., Degrapol®) [21,22,35]. Similarly, polymers have been synthesized by diesterification between L-phenylalanine and 1,4-cyclohexane dimethanol to yield a diester diamine. Soft segments used in these polymers were polycaprolactone diol (PCL) or polyethylene oxide (PEO) to vary the physicochemical and degradation characteristics of the polymers. These urethane elastomers exhibit variations in their physical properties, i.e., PEO-containing polymers are hydrophilic, are softer, and degrade faster than PCL-containing hydrophobic, hard, slow-degrading polymers [24]. While both of these polymers degrade in vitro, the PEO (MW 1000)/Phenylalanine polymers degrade much faster in vitro than PCL (MW 1250)/phenylalanine–based polymers. Polyesterurethanes synthesized with 4,4-methylene diphenyl-isocyanate (MDI) and poly(butylene adipate) diol and 1.4 butane diol have been shown to be biodegradable [61,62,73]. However, these polymers degrade slowly with an approximate decrease of 20% in molecular weight over a period of 8 weeks. Similarly, poly(urethane–urea) as segmented block polymers of MDI and poly(tetramethylene glycol) containing dehydropiandrosterone (DHEA) have been shown to degrade in aqueous solutions in vitro and postimplantation in vivo. In these polymers the rate of degradation is dependent upon the concentration of DHEA present in the polymer, i.e., the polymers containing 10% DHEA degrade faster than those containing 5% DHEA [74]. Urethanes synthesized with co-condensation of telechelic acid, low molecular mass poly[(R)-3-hydroxybutyric acid-co-(R)-3-hydroxyvaleric acid]-diol (PHB, forming crystalline domains) and poly[glycolide-co-(-caprolactone)]-diol (present in amorphous domains), coupled with aliphatic 2,2,4-trimethylhexamethylene diisocyanate (TMID) as a chain extender, are also biodegradable in vivo. However, these polymers are also slow degrading, taking in excess of 12 weeks for complete degradation in vivo. Interestingly, polymers synthesized with toluenediisocyanate (TDI) and poly[(R)-3hydroxybutyric acid-co-(R)-3-hydroxyvaleric acid]-diol (PHB/HV-diol) or polycaprolactone diol (PCL-diol) have been shown to exhibit biodegradability [36]. However, these polymers degrade very slowly, i.e., 50% in a year [71]. This degradation is likely due to the hydrolytic instability of polyester contents of the polymer. In response to an increased need for biodegradable polymers, use of polyhydroxyalkanenoates and polyhydroxybutyrate has been suggested. These materials are of interest due to their thermoplastic properties, biodegradability, and biocompatibility [75]. These polymers are likely to be useful as soft segments in biodegradable polyurethanes. VIII BIOCOMPATIBILITY OF POLYURETHANES One of the most important factors in the application of tissue engineered scaffolds is their biocompatibility. Biocompatibility of a material refers to its ability to be tolerated by the host immune system. Additionally, in materials that degrade or erode, the products released from the polymer must be nontoxic and nonimmunogenic. Because of their successful use in biomedical devices, such as vascular catheters, total artificial hearts, ventricle assist devices, and small caliber vascular grafts for arterial reconstruction [10,13,14,76], their biocompatibility has been extensively studied [56,68,77–79]. The host immune response to implanted materials is a natural reaction. The magnitude of this response varies with each material and depends upon its chemical structure and

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its recognition by the cells of the immune system. Macrophages and neutrophils are the primary immune cells involved in the initiation of the immune response. Biodurable and/or hydrolytically and oxidatively resistant polyurethanes activate both macrophages and neutrophils, and their responses to polyurethanes have been studied in vivo and in vitro [68,70,71,78,79]. Macrophages, but not neutrophils, adhere to bioresistant polyetherurethanes via complement C3 receptors and are activated on their surface as examined by changes in their morphology [68,79,80]. While macrophages do not produce tumor necrosis factor in presence of polyurethanes, the release of superoxide by macrophages and neutrophils is often observed in the presence of polyurethanes [81]. However, the activation of these cells appears to be mediated by adherent serum proteins, such as immunoglobulins, human albumin, -2 macroglobulin, complement factors C3b, thrombospondin, vitronectin, and von Willebrand factor [79]. These findings suggest that the polyurethanes themselves may not be the cause of immunoreactivity. Recent biocompatibility testing of biodegradable polyurethane scaffolds in experimental animals has shown marked promise for their application in tissue engineering. Biodegradable polyesterurethane foams, Degrapol®, synthesized with TDI and PHB/HVdiol or PCL-diol serve as a compatible substrate for chondrocytes and support chondrocytic adhesion, cell proliferation, and phenotype as assessed by collagen type II/I synthesis in vitro [22]. Degrapol also exhibits favorable tissue compatibility in vivo with 50- to 250-m thick capsule formation [36]. However, these polymers are slow degrading [21]. Chemical analysis of degradation products of TDI, polyesterdiol, and ethylenediamine-based polyurethanes by mass spectrometry has shown that toluene diamine is not liberated from these polymers, suggesting that biodegradable polyurethanes employing TDI as the hard segment may have less toxicity than believed earlier [62,73]. Several studies have shown that polyesterurethanes synthesized with LDI-based hard segments are biocompatible in vitro and in vivo [25,26,28,29,62,73]. Furthermore, incorporation of poly(L-lactide) or 50:50 poly(lactide-co-glycolide) in these matrices improves cell adhesion and growth and increases hydrophilicity of these polymers. Similarly, 4,4methylene diphenyl-diisocyanate and poly(butylene adipate)-diol–based biodegradable polyurethanes support osteoblast cell attachment in vitro. However, cell attachment is dependent upon polymer surface morphology, where cells adhered to a particulate surface better than a smooth surface of the polyurethane [73]. Recently, poly(urethane–urea) matrices with LDI as the hard segment and glucose, glycerol, or polyethylene glycol as soft segments have been shown to be nontoxic in vitro. All of these polymers support attachment and proliferation of various cell types, i.e., chondrocytes, bone marrow stromal cells, endothelial cells, and osteoblasts (Fig. 11). Additional testing of these polymers as subcutaneous implants in vivo has shown that LDI-glucose poly(urea–urethanes) are nontoxic and do not induce significant foreign body responses or antibody formation over a period of 8 weeks in vivo [28]. However, a thin capsule of a few cell layers does form around these polymers (Fig. 12). Since slower degrading polyurethanes described previously (Degrapol) exhibit thicker capsule formation than the faster degrading urethanes (LDI-glucose), it is possible that capsule formation may be a function of the prolonged contact time of the polymer with the host tissue. Biodegradable polyurethanes synthesized by esterification of phenylalanine and 1,4-cyclohexane dimethanol to yield a diester and polymerized with polycaprolactonediol and polyethylene oxide also exhibit biocompatibility in vitro [24].

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Figure 11 Morphology of bone marrow stromal cells (BMSC) grown on LDI-glycerol polyurethane network by scanning microscopy. (A) BMSCs showing spreading and adherence on LDIglycerol polymer 4 h after culture. (B) BMSCs grown on LDI-glycerol for 7 days showing cell spreading and attachment. (C) Cross-section of an LDI-glycerol foam showing growth of BMSCs in multiple layers in the pores of the foam after 4 weeks of culture.

IV SUMMARY While biodurable urethanes have been used in biomedical applications for decades, the use of biodegradable urethanes for tissue engineering purposes has been only recently proposed. Biodegradable urethanes synthesized with LDI or other amino acid–based hard segments/chain extenders and different diols or polyols degrade mainly by hydrolytic cleavage and liberate amino acids and polyols, such as glucose or polyesters. These degradation products appear to be nontoxic and are favorably tolerated by the host. Biodegradable polyurethanes offer many attributes that are advantageous for a scaffolding matrix used in tissue engineering applications, such as extensive structural, mechanical, and physical versatility, as well as an ability to bind and release signaling molecules. Further elucidation of interactions between living tissues and biodegradable urethanes is essential to enable the development of biodegradable implants with increased long-term safety and biocompati-

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Figure 12 A histological assessment of the biocompatibility of LDI-glucose polymer. Cross-sections of subcutaneous sites implanted with LDI-glucose matrix following: (A) 2 weeks of implantation showing granulation tissue and extensive vascularization around the polymer; (B) 4 weeks showing a thin capsule formation around the polymer without apparent toxicity to the tissue. (CT indicates connective tissue.)

bility. When developed to their full potential, these broadly diversified polymers may become widely used scaffolding matrices for tissue engineering applications. ACKNOWLEDGMENTS The authors wish to acknowledge the support of this work from Central Medical Research Fund, University of Pittsburgh, and Pittsburgh Tissue Engineering Initiative. REFERENCES 1. Ishaug-Riley S. L., Crane G. M., Gurlek A., Miller M. J., Yasko A. W., Yaszemski M. J., Mikos A. G. 1997. Ectopic bone formation by marrow stromal osteoblast transplantation using poly (DL-lactic-co-glycolic acid) foams implanted into the rat mesentery. J. Biomed. Mater. Res. 36:1–8. 2. Ishuang S. L., Crane G. M., Miller M. J., Yasko A. W. 1997. Bone formation by three-dimensional stromal osteoblast culture in biodegradable polymer scaffolds. J. Biomed. Mater. Res. 36:17–28.

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8 Significance of Drug Delivery in Tissue Engineering Yoshito Ikada Suzuka University of Medical Science, Suzuka, Japan Yasuhiko Tabata Institute for Frontier Medical Sciences, Kyoto University, Kyoto, Japan

I INTRODUCTION Tissue engineering is one of the biomedical technologies that assists the regeneration of body tissue by promoting to the greatest extent possible self-repair or substitution of the biological functions of damaged organs by using cells. For successful tissue regeneration, it is key to artificially create a site suitable for regeneration induction at the defect. This can be achieved only by taking advantage of an artificial scaffold with three-demensional structures for cell proliferation and differentiation as well as growth factors. Growth factor is often required to promote tissue regeneration and can induce angiogenesis which is effective in supplying oxygen and nutrients to survive the cells transplanted for organ substitution. However, one cannot always expect the biological effects of growth factor because of its poor in vivo stability, unless its delivery system is contrived. This chapter describes recent experimental data emphasizing the significance of drug delivery in tissue engineering and gives a brief overview of biodegradable polymers used for this purpose. II WHAT IS TISSUE ENGINEERING? When body tissue or organ is severely injured, largely lost, or functionally wrong, it is clinically treated with either reconstruction surgery or organ transplantation. Reconstruction surgery is accomplished with biomedical materials or devices. Biomedical devices made from manmade materials alone cannot perform all of the functions of a single organ and 145

Figure 1 Basic principle in tissue engineering.

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therefore cannot prevent progressive patient deterioration. Although these therapies have saved and improved countless lives of patients, they remain imperfect solutions. One of the largest problems for organ transplantation is the lack of donor tissues or organs. Additionally the permanent medication of immunosuppessive agents causes various side effects. One promising approach to tackle these problems is to enable the self-healing potential of patients to regenerate their own body tissues and organs. A biomedical field using this new therapeutic approach is tissue engineering. The objective of tissue engineering is to regenerate natural tissues as well as to create biological substitutes for defective or lost tissues and organs by making use of cells. III FACTORS NECESSARY FOR TISSUE ENGINEERING How can a defective or lost tissue be regenerated? It is undoubtedly necessary for tissue regeneration to increase the number of cells constituting the tissue as well as to reconstruct a structure to support the proliferation and differentiation of cells, the so-called extracellular matrix (ECM). In addition, growth factors are often required to promote tissue regeneration, depending on the tissue type. The mechanism on tissue regeneration has been currently elucidated at the cellular and genetic levels and found to be fundamentally similar between newt or hydra and human. In either case, the stem cells of multipotential and proliferation ability play a key role in tissue regeneration. The factors necessary for tissue engineering include cells, the scaffolds for cell proliferation and differentiation, and growth factors. The cells currently used in tissue regeneration do not always need stem cells and it is possible to utilize precursor or blastic cells, defined as a cell stage which is intermediate between stem cells and maturing cells. It would be ideal if we could obtain cells from patients themselves, but cell harvest from individual patients is usually not easy. In such cases, heterogenous cells will be used, although care should be taken for immunoisolation. Embryonic stem (ES) cells are also of high potentiality, but the ethical issue should be carefully considered for their clinical usage. It is well known that the ECM is not only a physical support of cells, but also has an important influence on cell proliferation and differentiation or morphogenesis, which contributes to tissue regeneration and organogenesis. It is unlikely that a large-sized defect of tissue will be naturally regenerated and repaired only by supplying cells to the defect. One possible way to facillitate this process is to provide a site suitable for induction of tissue regeneration at the defect by placing a scaffold in advance as an artificial ECM which supports cell attachment and the subsequent proliferation and differentiation. It is highly expected that self-derived cells residing around the scaffold or cells preseeded in the scaffold proliferate and differentiate on the foundation of the provided scaffold if the artificial ECM is compatible to the cells. Once a new tissue is regenerated, the tissue eventually produces the intact ECM. The scaffold remaining sometimes causes physical hindrance during the process of tissue regeneration. The third factor is growth factor. However, the direct injection of growth factor in the solution form into the regeneration site is generally not effective, as the injected growth factor is rapidly diffused out from the site. To enable the growth factor to efficiently exert the biological effects, the techniques of drug delivery are practically available. One promising technique is the controlled release of growth factor at the site of action over an extended time period by incorporating the factor into an appropriate carrier. It is also highly possible that the growth factor is protected against its proteolysis, at least as far as it is incorporated in the release carrier, for prolonged retention of the activity in vivo. The release carrier should be degraded in the body since it is not needed any more after the

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growth factor release is completed. Thus, it is a key of tissue engineering to create an environment suitable for induction of tissue regeneration by the scaffold and drug delivery technologies. The tissue engineering will never be achieved without this creation for the regeneration environment. Many materials have been investigated with regard to their medical and pharmaceutical applications. Considering the scaffold and drug delivery applications described, it is undoubtedly preferable to take advantage of biodegradable materials for tissue engineering. IV BIODEGRADABLE MATERIALS FOR TISSUE ENGINEERING Table 1 summarizes synthetic and natural polymers of biodegradable nature. The synthetic biodegradable polymers used clinically are the homopolymer of lactide and its copolymers with glycolide. Their biodegradable pattern can be controlled by changing the molecular weight and copolymer compositions, which is described in another chapter of this book in more detail. A poly(anhydrides) and a poly(cyanoacrylate) are being used as carriers of an antitumor agent and a surgical adhesive, respectively. Other synthetic polymers have been experimentally investigated aiming at their biomedical and pharmaceutical applications. Among natural polymers, proteins (collagen, gelatin, fibrin, and albumin) and polysaccharides (chitin, hyaluronic acid, cellulose, and dextran) have been medically and pharmaceutically employed. Generally the degradation of polymers is driven by hydrolytic and enzymatic cleavage of their main chain. Most synthetic polymers are fundamentally degraded by hydrolysis, although poly(amino acids) show enzymatic degradation. On the contrary, the natural polymers are all degraded enzymatically. There is a degradation manner in which a polymer became water soluble with the chemical elimination of the side chain, disappearing from the implanted site. No metal materials degradable in the body are present to be corroded. There are some biodegradable ceramics, for example, tricalcium phosphate and calcium carbonate. Hydroxyapatite is not practically degraded in the body because of its extremely low solubility in water. V CLASSIFICATION OF TISSUE ENGINEERING Tissue engineering is classified into two categories according to where it is performed: in vitro and in vivo tissue engineering (Table 1). In vitro tissue engineering involves tissue regeneration and organ substitution (bioartificial organs). A In Vitro Tissue Engineering If a tissue can be reconstructed in vitro in factories or laboratories on a large scale, we can supply the tissue construct to patients when it is needed. This is an ideal therapy and available for commercialization. However, for in vitro tissue engineering, it is highly important to arrange an environment required for tissue reconstruction by providing all the essential materials. This contrasts with in vivo tissue engineering, where most of the materials necessary for tissue regeneration are automatically supplied by the host living body. Therefore, reconstruction of only a few tissues have been attempted in vitro: skin (dermis and epidermis), articular cartilages, and arteries. An artificial skin composed of human dermal and epidermal layers has been prepared in vitro [1]. This is applicable not only to the skin regeneration but also to the skin analog for drug testing. Another application of the in vitro tissue engineering is substitution of organ functions by the use of allo- or xenogeneic cells. Such engineered organs are called bioartificial

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Table 1 Synthetic and Natural Polymers of Biodegradable Nature Based on the Chemical Structure Synthetic polymer The name and structure of chemical bond Ester MCMOM  O

Anhydride MCMOMCM   O O

Example

Natural polymer (animals, plants, microbes) The name and structure of chemical bond

Polylactide, Ester polyglycolide, MCMOM  lactide–glycolide O copolymer, poly(caprolactone), poly(p-dioxane), poly(-malic acid) Poly(anhydrides)

Ortho ester

Poly(ortho esters)

Carbonate MOMCMOM  O

Polycarbonates

Phosphazene MNBPM

Poly(phosphazenes)

Peptide MNHMCM  O

Poly(amino acids)

Peptide (protein) MNHMCM  O

Phosphoric ester O  MPMOM  O

Poly(phosphoric ester–urethanes)

Phosphoric ester (nucleic acid) O  MPMOM  O

Carbon–carbon CN  MCH2MCM 

Poly(cyanoacrylates)

Glycoside (polysaccharide)

Example Poly(-hydroxybutyrate), poly(malic acid)

Chitin, chitosan, hyaluronic acid, pectin, pectic acid, galactan, starch, dextran, pullulan, agarose, heparin, alginate, chondroitin6-sulfate Collagen, gelatin, fibrin, albumin, gluten, polypeptides, elastin, fibroin, enzyme Deoxyribonucleic acid (DNA), ribonucleic acid (RNA)

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organs because they are composed of heterogeneic cells and manmade membranes or porous constructs for immunoisolation to protect the cells from host attack and maintain the cell function. Liver and pancreas are target organs that have attracted much attention of researchers on bioartificial organs [2]. B In Vivo Tissue Engineering Most of the current tissue engineering is performed in vivo with or without biodegradable scaffolds. If the healthy ECM is still available in the body, no artificial scaffold is needed. Following intravenous injection into patients with leukemia, the hemotopoietic stem cells isolated from the human blood can differentiate into final functional cells in the bone marrow (blood cell transplantation). Eye-related stem cells are being used for regeneration of defective cornea and retina [3], while the transplantation of myocardial cells has been tried for myocardial infarction therapy. For the regeneration of a large-sized defect, it is necessary to use a biodegradable scaffold. The scaffold is implanted with or without cell seeding. One of the most popular scaffolds in tissue engineering is a collagen sponge [4]. The sponge functions as a biodegradable scaffold for the proliferation of cells originated from the surrounding healthy tissue and the subsequent secretion of natural collagen. The collagen scaffold should degrade in the body so it will not be a physical hindrance to the newly regenerated tissue. Tissue engineering approaches using a collagen sponge or a biodegradable polymer sheet with no cell seeding have been attempted for regeneration of the skin dermis [5], trachea [6], esophagus [7], and dura mater [8]. Most body tissues are not able to regenerate unless the scaffold used is seeded with the corresponding cells. Regeneration of epidermis and cartilage necessitates seeding of the scaffold with keratinocytes and chondrocytes, respectively. Combination of cells isolated from a blood vessel and small intestine with a biodegradable scaffold achieved the in vivo regeneration of the respective organs [9,10]. Bone marrow cells are also available. Bone regeneration readily takes place by using them together with a scaffold [11]. This is because bone marrow cells contain mesenchymal stem cells that can differentiate into the osteocytic lineage [12]. It is possible to seed more than one cell for regeneration of tissue composed of several subtissues. For example, phalanges and small fingers could be reconstructed by seeding three different scaffolds with periosteum, chondrocytes, and tenocytes for bone, cartilage, and tendon (ligament), respectively [13]. There are some cases in which growth factor is required for in vivo tissue engineering. The type of growth factor depends on the tissue and on the site where the tissue is produced. Further addition of growth factors to the scaffold seeded with cells will accelerate the tissue regeneration. Some concrete examples of this approach will be described later. When a body defect is generated, the defect space will be soon filled with the fibrous tissue produced by fibroblasts, which are ubiquitously present in the body and can rapidly proliferate. Once this ingrowth of fibrous connective tissue into the space takes place, more tissue repair or regeneration cannot be expected. To prevent the tissue ingrowth, we need biomedical materials called barrier membranes. The objective of membranes is to make space for tissue regeneration and prevent the undesirable tissue ingrowth, permiting repair of the defective tissue. The examples include guided channel for broken peripheral nerve fibers [14] and guided regeneration of lost periodontal tissues and alveolar bone [15]. The membrane used for such a case should be prepared from biodegradable materials (Table 1) because it is no longer needed after completion of tissue regeneration.

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VI TISSUE ENGINEERING BY GROWTH FACTOR RELEASE If the tissue to be repaired has a high activity toward regeneration, a new tissue will be formed in the biodegradable scaffold matrix by active, immature cells infiltrated from the surrounding healthy tissue. However, additional means are required if the regeneration potential of the tissue is very low because of, for instance, low concentration of the cells and growth factors responsible for new tissue generation. The simplest method to counter this is probably to supply growth factors to the site of regeneration for cell differentiation and proliferation. As described, it is necessary for growth factors to use the controlled release system. Table 2 summarizes research on tissue regeneration through the combination of growth factors with the various carriers. All the research without exception indicates that combination with a carrier is necessary to induce growth factor function in vivo for tissue regeneration. However, in spite of this necessity, little investigation has been done on growth factor release from the carrier and the release effect on tissue regeneration. There are some reports that usage of growth factor in solution form was effective in enhancing tissue regeneration. However, the dosage is too high to accept therapeutic clinical trials. Although not included in Table 2, the gene encoding growth factor has recently been applied to tissue engineering [66]. If such a gene is transfected into the cells existing in the site of regeneration, it is possible that the cells secrete the growth factor for a certain time period, resulting in promoted tissue regeneration. Therapy for ischemic diseases [67] and bone tissue regeneration [68] has been achieved by growth factor genes. VII CONTROLLED RELEASE OF GROWTH FACTOR FROM BIODEGRADABLE HYDROGEL One of the biggest problems in protein release technology is the loss of biological activity of the protein released from a protein–polymer formulation. It has been demonstrated that this activity loss results from denaturation and deactivation of protein during the formulation process with a polymer matrix. Therefore, a new formulation method of release carrier with polymers should be exploited to minimize protein denaturation. From this viewpoint, polymer hydrogel may be a preferable candidate as protein release carrier because of its biosafty and its high inertness toward protein drugs. However, it will be impossible to achieve the controlled release of protein over a long time period from hydrogels since the rate of protein release is generally diffusion controlled through aqueous channels inside the hydrogels. Thus, one possible practical approach is to immobilize a growth factor in a biodegradable hydrogel, allowing the immobilized factor to release as a result of hydrogel biodegradation. In such a release system, the growth factor release can be controlled only by changing the in vivo hydrogel degradation. Chemical and physical methods are available to immobilize growth factors into the hydrogel. Since the former method often results in the denaturation and activity lowering of protein, it is preferable to follow the latter one in terms of growth factor efficacy. Actually such physical immobilization can be observed for growth factors existing in the body [69]. Since it is known that generally some growth factors possess a positively charged site on the molecular surface, they are normally stored in the body, being ionically complexed with acidic polysaccharides of the ECM, such as heparan sulfate and heparin. This complexation protects growth factors from denaturation and enzymatic degradation in vivo. The growth factor is released form the ECM complex as a result of polysaccharide degradation according to the need.

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Table 2 Experimental Trials for Tissue Regeneration by Combination of Growth Factor with the Carrier Growth factor BMP

Carrier

Animal

PLA Collagen sponge

Dog Rat Dog, monkey

rhBMP-2

-TCP Porous HA Porous PLA

rhBMP-7

PLA microsphere Collagen sponge Gelatin PLA-coated gel gelatin sponge Porous HA PLA-PEG copolymer Collagen

Rabbit Rabbit Dog Rat Rabbit Dog Rabbit Dog, monkey

EGF aFGF bFGF

NGF TGF-1

Agarose PVA PVA Alginate Alginate Agarose/heparin Amylopectin Gelatin

Fibrin gel Collagen minipellet Collagen Poly(ethylene-covinyl acetate) Collagen minipellet PLGA PEG Gelatin Plaster of Paris, PLGA TCP Porous HA Collagen

Monkey Rat Dog Dog Hamster Rat Mouse Mouse Mouse Mouse, pig Mouse Mouse

Tissue regenerated Long bone Long bone Periodontal ligament and cementum Long bone Skull bone Spinal bone Skull bone Skull bone Periodontium Skull bone Long bone, jaw bone, skull bone Skull bone Long bone

Reference 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 34 35, 36 37 38, 39, 40

Rabbit, monkey Dog Mouse Rabbit Mouse Rat

Spinal bone Long bone Angiogenesis Dermis Angiogenesis Angiogenesis Angiogenesis Angiogenesis Angiogenesis Angiogenesis, dermis, adipogenesis Skull bone Nerve Angiogenesis Long bone Cartilage Nerve

Rabbit Rat Rat Rabbit Rat

Nerve Nerve Dermis Skull bone Skull bone

48 49 50 51 52

Dog Dog Baboon Mouse

Long bone Long bone Skull bone Dermis

53 54 55 56

41, 42 43 44 45 46 47

continued

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Table 2 Continued Growth factor PDGF-BB

VEGF HGF IGF-I IGF-I/bFGF PDGF/IGF-I

Carrier Porous HA Collagen Chitosan Collagen Alginate Gelatin PLGA-PEG PLGA-PEG Titanium implant

Animal Rabbit Rat Rat Mouse Mouse Mouse Rat Rat Dog

Tissue regenerated Long bone Dermis Periodontal bone Angiogenesis Angiogenesis Angiogenesis Adipogenesis Adipogenesis Jaw bone

Reference 57 58 59 60 61 62 63 64 65

aFGF, acid fibroblast growth factor; bFGF, basic fibroblast growth factor; BMP, bone morphogenetic protein; rHBMP, recombinant human bone morphogenetic protein; EGF, epidermal growth factor; HGF, hepatocyte growth factor; HA, hydroxyapatite; IGF-I, insulin-like growth factor-I; NGF, nerve growth factor; PDGF-BB, platelet-derived growth factor–BB; PEG, poly(ethylene glycol); PLA, polylactide; PLGA, glycolide-lactide copolymer; PVA, poly(vinyl alcohol); TCP, tricalcium phosphate; TGF, transforming growth factor; VEGF, vascular endothelial growth factor.

We have created a release system of growth factors which mimics that in the living body. Figure 2 shows an idea on the controlled release of growth factor from biodegradable hydrogel based on intermolecular interaction forces. For example, a hydrogel is prepared from a biodegradable polymer with negative charges. The growth factor with a positively charged site is electrostatically interacted with the polymer chain to allow physical immobilization in the hydrogel carrier. If an environmental change, such as increased ionic strength, occurs, the immobilized growth factor will be released from the factor–carrier formulation. Even if such an environmental change does not take place, degradation of the carrier itself will also lead to growth factor release. Because the latter is more likely to happen

Figure 2 Conceptual illustration of growth factor release from biodegradable hydrogel based on physical interaction forces.

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in vivo than the former, the release carrier is preferentially prepared from biodegradable polymers. For growth factor release on the basis of the physical interaction forces, it is absolutely necessary to employ a biosafe polymer with groups able to interact as the carrier material. In addition, if biodegradability is required, the material to be used will be restricted to natural polymers with charged groups. Therefore, we have selected gelatin, which has been extensively used for industrial, pharmaceutical, and medical purposes. The biosafety of gelatin has been proved through its long clinical usage. Another unique advantage is the electrical nature of gelatin which changes according to the collagen processing method. For example, the alkaline process through hydrolysis of amide groups of collagen yields gelatin having a high density of carboxyl groups, which makes the gelatin negatively charged. If a growth factor to be released has the positively charged site in the molecule which interacts with acidic polysaccharides present in the ECM, the negatively chraged gelatin of “acidic” type is preferable as the carrier material. It was found that as expected, basic fibroblast growth factor (bFGF), transforming growth factor 1 (TGF-1), hepatocyte growth factor (HGF), or platelet-derived growth factor (PDGF) was sorbed into the acidic gelatin hydrogel mainly due to the electrostatic interaction [70]. Animal experiments revealed that the hydrogels prepared from the acidic gelatin were degraded in the body [71]. The degradation period of hydrogels depends on their water content, which is a measure of crosslinking extent: the higher the water content of the hydrogels, the faster their in vivo degradation. When traced in the back subcutis of mice by use of gelatin hydrogels incorporating 125I-labeled bFGF or 125I-labeled gelatin hydrogels incorporating bFGF, both the residual radioactivity decreased with implantation time and the decrement rate increased with the decreased water content of hydrogels. The time profile of bFGF radioactivity remaining in the hydrogel depended on the hydrogel degradability and was in good accordance with that of hydrogel radioactivity remaining [71]. These findings strongly indicate that the growth factor release is governed mainly by hydrogel degradation, as described in Fig. 2. As a result, the release period is not influenced by the hydrogel shape, but controllable only by changing the degradation rate of hydrogel [72]. It should be noted that gelatin hydrogels can be formulated into different shapes of discs, tubes, sheets, granules, and microspheres [72,73]. VIII TISSUE ENGINEERING BY GELATIN HYDROGELS INCORPORATING GROWTH FACTOR As described in the previous section, the gelatin hydrogel was found to be a good carrier for the controlled release of growth factor. In this section, several experimental results on angiogenesis, bone regeneration, and adipogenesis achieved by this release system are described to emphasize the significance of drug delivery in tissue engineering. A Angiogenesis Basic fibroblast growth factor was originally characterized in vitro as a growth factor for fibroblasts and capillary endothelial cells and in vivo as a potent mitogen and chemoattractant for a wide range of cells. In addition, bFGF is reported to have a variety of biological activities [74] and be effective in enhancing wound healing through induction of angiogenesis and regeneration of bone, cartilage, and nerve. The gelatin hydrogel was effective in enhancing the in vivo angiogenic effect of bFGF. When gelatin hydrogels in-

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corporating bFGF were subcutaneously implanted into the mouse back, angiogenic effect was observed around the implanted site, in marked contrast to the site implanted with bFGF-free, empty gelatin hydrogels or injected with aqueous solution of bFGF [70]. No angiogenesis was induced by the injection of bFGF solution even when the dose was increased to 1 mg per site. This must be due to a rapid elimination of bFGF from the injection site [75]. On the contrary, the gelatin hydrogel incorporating bFGF induced significant angiogenesis at doses as low as 30 g per site. As shown in Fig. 3, the maintenance period of hydrogel-induced angiogenic effect could be changed by the hydrogel water content and prolonged as the water content became lower [75]. It is likely that the hydrogels with lower water contents are degraded and consequently release bFGF of biological activity in vivo more slowly than those with higher water contents, leading to a prolonged angiogenic effect. Similar enhanced and prolonged angiogenic effect was also observed upon using gelatin hydrogel incorporating bFGF of microsphere type [72]. The technique to induce in vivo angiogenesis is very important for tissue engineering. The two objectives of angiogenesis induction include the therapy of ischemic disease and angiogenesis in advance for cell transplantation. Here, as in the former example, the therapy of ischemic myocardium by gelatin hydrogels incorporating bFGF is shown. When injected into the myocardial infarction which was prepared by ligating the left anterior descending (LAD) coronary artery of dog heart, the gelatin microspheres incorporating bFGF induced regeneration of collateral coronary arteries as well as improved the motion of myocardium in the ischemic region (Fig. 4). Neither therapeutic effect was observed with the injection of bFGF solution at the same dose. It should be noted that sufficient supply of nutrients and oxygen to cells transplanted in the body is indispensable for cell survival and the maintenance of biological functions. Without sufficient supply, cells preseeded in a scaf-

Figure 3 The time course of angiogenesis induced by gelatin hydrogels incorporating bFGF with a water content of gelatin hydrogels: 98.0 (clear circle); 94.1 (filled circle); 90.4 (clear triangle); and 85.5 wt% (filled triangle); a noncrosslinked gelatin hydrogel incorporating bFGF (square). The bFGF dose was 100 g per mouse. The dotted line indicates the weight of tissue hemoglobin in the corresponding area of untreated, normal mice. The asterisks indicate significance at p0.05 against the value of control mice at the corresponding day.

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Figure 4 Left coronary angiograms of ischemic dog heart 1 week after intramyocardial injection of (a) bFGF solution and (b) gelatin microspheres incorporating bFGF. bFGF was bilaterally injected at the distal side of the LAD ligated portion (indicated by the mark II) at the dose of 100 g per heart. The hydrogel water content was 95.0 wt%.

fold for tissue regeneration would hardly survive following implantation of the scaffold into the body. Such a situation must be met for allo- or xenogeneic cells transplanted into the body for organ substitution. For successful cell transplantation, the nutrients and oxygen supply is a significant problem, comparable to immunoisolation. In both the cases, a promising way is to induce angiogenesis throughout the transplanted site of cells by making use of angiogenic growth factors. After implantation into the subcutaneous tissue of streptozotocin-induced diabetic rats, the allogeneic pancreatic islet encapsulated by a poly(vinyl alcohol) (PVA) bag effective for immunoisolation did not always contribute to the normalization of blood glucose level for a long time. This was due to the islet death caused by poor blood supply in the subcutaneous tissue compared with that in the peritoneal cavity. However, advance angiogenesis at the site of cell transplantation induced by gelatin microspheres incorporating bFGF enabled the encapsulated islet to improve the survival rate, resulting in the prolonged maintenance period of normalized glucose level in the blood (Fig. 5). This angiogenic effect on the prolonged cell survival was observed for hepatocyte transplantation. A biodegradable sponge of L-lactic acid and -caprolactone copolymer was implanted into the peritoneal cavity of rats followed by gelatin microspheres incorporating bFGF placed into the sponge for advance induction of angiogenesis. After injection into the sponge 1 week after bFGF-induced angiogenesis, allogeneic rat hepatocytes remained alive in the sponge even 2 months later. In the case of the control sponge without gelatin microspheres incorporating bFGF, the transplanted hepatocytes did not survive for such a long time period [76]. These findings clearly indicate that the advance induction of angiogenesis at the site of cell transplantation was effective in achieving successful cell transplantation. B Bone Regeneration Gelatin hydrogels incorporating bFGF were found to have a promising potential for bone repair [41,42]. For example, when implanted into a monkey skull defect, the gelatin hydrogel incorporating bFGF promoted bone regeneration at the defect and closed the defect 21 weeks after implantation. On the contrary, use of bFGF-free, empty gelatin hydrogels and the same dose of bFGF in the solution form resulted in no bone regeneration, while remarkable ingrowth of soft connective tissues was observed in the bone defect (Fig. 6). Mea-

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Figure 5 The time course of blood glucose level of diabetic rats following subcutaneous implantation of allogeneic pancreatic islet encapsulated by a poly(vinyl alcohol) hydrogel bag. The bFGF dose was 100 g per site and the hydrogel water content was 95.0 wt%.

Figure 6 (A) Histological cross-sections around skull defects of monkey 21 weeks after treatment with PBS, (B) an empty gelatin hydrogel, (C) free bFGF, and (D) a gelatin hydrogel incorporating bFGF: b, bone; d, dura mater; c, connective tissue; nb, new bone. (A–D: HE staining, 40). The bFGF dose was 100 g per defect and the hydrogel water content was 85.0 wt%.

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surement of the bone mineral density (BMD) at the skull defect of monkeys revealed that gelatin hydrogels incorporating bFGF enhanced the BMD to a significantly higher extent than free bFGF, irrespective of the hydrogel water content. The BMD of empty gelatin hydrogels was similar to that of an untreated group, indicating that the hydrogel presence did not disturb bone healing at the defect. In a histological study, the hydrogel implantation increased the number of osteoblasts residing near the bone edge of defect and retained it at a significantly high level over the time range studied. The retention period of the enhanced cell number became longer with a decrease in the hydrogel water content. These findings indicate that controlled release from the gelatin hydrogel enabled bFGF to activate osteoblasts for an extended time period, resulting in induced regeneration of skull bone. It is known that both TGF-1 and bone morphogenetic protein (BMP) also promote bone regeneration [25,51,77,78]. We have succeeded in bone repairing at the skull defects of rabbits and monkeys by the controlled release of TGF-1 from the gelatin hydrogel, in marked contrast to free TGF-1 even at higher doses [51]. However, the repairing effect depended on the water content of hydrogels and was reduced with the increased or decreased hydrogel water content. It is possible that too rapid a degradation of the hydrogel causes a short period of bFGF release, resulting in no induced bone regeneration. Conversely, the long-term retention of hydrogels due to their slow degradation would physically hinder bone regeneration. As a result, it is likely that the hydrogel with an optimal biodegradability induced complete bone regeneration at the skull defect [51]. In this case, the hydrogel functions as the carrier of growth factor as well as the barrier to prevent the ingrowth fibrous tissues into the bone defect. Balance of the time course between the two hydrogel functions would result in better bone repair. We have recently succeeded in the controlled release of BMP-2 by hydrogels of a gelatin type in which the time period of BMP-2 release can be regulated by changing the hydrogel biodegradation. This controlled release system enabled BMP-2 to induce formation of bone tissue ectopically or orthotopically at doses lower than with the application of BMP-2 solution. There are some cases in which the controlled release of growth factor alone does not always achieve bone repair at a large-sized defect. For one trial, we have utilized cells with osteogeneic potentials and combined them with the release system of growth factor. Among them are mesenchymal stem cells (MSC) present in the bone marrow. We have demonstrated that application of combined MSC and gelatin microspheres incorporating TGF-1 allowed complete closure of a large-sized defect of rabbit skulls by newly formed bone tissue, in marked contrast to that of either material alone [79]. C Adipogenesis When gelatin microspheres incorporating bFGF were mixed with a basement membrane extract (Matrigel) and subcutaneously implanted into the mouse back, de novo formation of adipose tissue was observed at the implanted site [40]. A histological study revealed that matured adipocytes were observed in the tissue mass newly formed following the implantation of the mixed microspheres and Matrigel, in contrast to that of either material alone. It is possible that the controlled release of bFGF induced angiogenesis, resulting in efficient proliferation and maturation of adipose precursor cells migrated in the angiogenesis-induced Matrigel, in addition to the direct biological actions of bFGF on precursor cells. Recently we succeeded in regeneration of adipose tissue by the preadipocytes isolated from fat tissues, gelatin microspheres incorporating bFGF, and a collagen sponge (Fig. 7). When the preadipocytes and the microspheres were placed into the collagen sponge and im-

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Figure 7 De novo formation of adipose tissue in the mouse subcutis 6 weeks after implantation of a collagen sponge including the mixture of preadipocytes and gelatin microspheres incorporating bFGF: (A) a collagen sponge including the mixture of preadipocytes and gelatin microspheres incorporating bFGF; (B) a collagen sponge including the mixture of preadipocytes and free bFGF; (C) a collagen sponge including preadipocytes; (D) a mixture of preadipocytes and gelatin microspheres incorporating bFGF; and (E) a collagen sponge including gelatin microspheres incorporating bFGF. (100, Sudan III staining). The bFGF dose was 10 g per site and the hydrogel water content was 95.0 wt%. Bar  300 m. planted into the back subcutis of mice, de novo formation of adipose tissue was observed at the implanted sponge site. Combination of all three materials was required for this adipogenesis. It is likely that the released bFGF increases the number of preadipocytes and the rate of differentiation into matured adipocytes in the collagen sponge as a scaffold, resulting in achievement of de novo adipogenesis. IX CONCLUDING REMARKS For regeneration of body tissues, a variety of growth factors act on cells forming a complex network, and the timing, site, and concentration of action are delicately regulated in the body. It is likely that the mechanisms of living systems will be clarified with rapidly advancing progress in cell biology, molecular biology, and embryology. Even so, it will be impossible to imitate the living systems with only the technologies currently available. However, this clarification will help us to understand which growth factor is key to achieve the regeneration of target tissue. If such a key growth factor is supplied to the necessary site at a suitable time period and concentration, we believe that the living body system will be naturally directed toward the process of tissue regeneration. Once the right direction is given, the intact system of the body will start to function, resulting in automatic achievement of tissue regeneration. It is no doubt that as far as growth factors are used in vivo, their controlled release is an essential technology. However, the present technology of controlled release is still poor to completely regulate the amount and period of growth factor to be released. Additionally, there is also the case in which it is not always enough only to control the amount and period of growth factor released. Therefore, at the present time, one possible practical way is to release only a certain growth factor necessary to increase the num-

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ber of precursor, blastic, or stem cells in vivo. It should be noted that it is practically impossible to control cell differentiation by the currently available technology of growth factor release. As described, tissue engineering will be the third choice of therapeutic medicine, equal to reconstructive surgery and organ transplantation. To enable the application of tissue engineering to clinical medicine, substantial collaborative research among materials, pharmaceutical, biological, and medical sciences and clinical medicine is needed to reach academic and technical maturity in tissue engineering. However, little research has been reported on the research of drug delivery aimed at tissue regeneration as well as organ substitution. Probably, one of the reasons is poor availability of growth factors on a large scale together with their high cost. There are technologies of drug delivery other than controlled release: prolongation of drug lifetime, improvement of drug absorption, and drug targeting. For example, a promising way to promote tissue regeneration is by making use of the technology to target a growth factor to the site to be regenerated with prolonging the in vivo lifetime. The technology of drug delivery will also be applicable to create the nonviral vector for gene transfection to prepare gene-engineered cells for tissue engineering. As tissue engineering is still in its infancy, it will take much more time to come into full bloom. This chapter underscores the increasing significance of drug delivery in the future progress of tissue engineering, and we hope it will induce interest in this research field. ACKNOWLEDGMENTS Our research projects were supported by a grant from the Research for the Future program by the Japan Society for the Promotion of Science (JSPS-RFTF96I00203). REFERENCES 1. Suzuki S., Matsuda K., Maruguchi T., et al. 1995. Further application of bilayer artificial skin. Br. J. Plast. Surg. 48:222–229. 2. Prokop A., Hunkeler D., Cherrington A. D. 1997. Bioartificial organs, sciences, medicine, and technologies. Ann. NY Acad. Sci. 831:249–298. 3. Tsubota K., Satake Y., Kaido M., et al. 1999. Treatment of severe ocular-surface disorders with corneal epithelial stem-cell transplantation. N. Eng. J. Med. 340:1697–1703. 4. Shimizu Y. 1998. Tissue engineering for soft tissue. In: The Tissue Engineering for Therapeutic Use, Ikada Y., Enomoto S., Eds. Elsevier Science: Amsterdam, pp. 119–122. 5. Yannas I. V., Burke, J. F. 1980. Design of an artificial skin. 1. Basic design principle. J. Biomed. Mater. Res. 14:65–81. 6. Okumura N., Nakamura T., Shimizu Y., et al. 1994. Experimental study on a new tracheal prosthesis made from collagen-conjugated mesh. J. Thorac. Cardiovasc. Surg. 108:337–341. 7. Takimoto Y., Nakamura T., Yamamoto Y., et al. 1998. The experimental replacement of a central esophageal segment with an artificial prosthesis with the use of collagne matrix and a silicone stent. J. Thorac. Cardiovasc. Surg. 116:98–106. 8. Yamada K., Miyamoto S., Nagata I., et al. 1997. Development of a dural substitute from synthetic bioabsorbable polymers. J. Neurosurg. 86:1012–1017. 9. Shinoka T., Shun-Tim D., Ma P. X., et al. 1998. Creation of viable pulmonary artery autografts through tissue engineering. J. Thorac. Cardiovasc. Surg. 115:536–546. 10. Kaihara S., Kim S. S., Kim B. S., et al. 2000. Long-term follow-up of tissue-engineered intestine after anastomosis to native small bowel. Transplantation 69:1927–1932. 11. Ohgushi H., Caplan A. I. 1999. Stem cell technology and bioceramics: from cell to gene engineering. J. Biomed. Mater. Res. (Appl. Biomater.) 48:913–927.

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9 Electrospinning of Polymer Scaffolds for Tissue Engineering Gary L. Bowlin, Kristin J. Pawlowski, Eugene D. Boland, David G. Simpson, John B. Fenn, and Gary E. Wnek Virginia Commonwealth University, Richmond, Virginia Joel D. Stitzel Virginia Tech, Blacksburg, Virginia

I INTRODUCTION Since the inception of the field of tissue engineering there has been a considerable effort to develop an “ideal” tissue engineering scaffold [1]. To date, investigators have developed materials such as collagen, poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and polycaprolactone (PCL) for use in matrix construction with useful, yet frequently unacceptable, clinical results [2,3]. The construction of an ideal scaffold requires multiple criteria to be met. The first is that the material must be biocompatible and function without interrupting other physiological processes. This functionality includes an ability to promote normal cell growth and differentiation while maintaining three-dimensional orientation/space for the cells. Second, the scaffold should not induce any adverse tissue reaction [3]. Also, scaffold production should be simple yet versatile [1,4–6] and should involve processes by which the scaffold can be easily reproduced to a wide range of shapes and sizes [3]. Once implemented in vitro or in vivo, the material should be completely resorbable, leaving only native tissue. Furthermore, we suggest that the proper in vivo phenotype cannot consistently be achieved if cells are presented with fibers with diameters equal to the cell size or, in many cases, an order of magnitude greater than the cell size. Observational evidence for this conclusion comes from a consideration of the native extracellular matrices of various tissues. Cells appear to have a highly intricate three-dimensional relationship with their extracellular structures. These structures, which are primarily composed of collagens, exhibit 165

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varying fiber diameters that are quite frequently one or more orders of magnitude smaller than the cell itself [7]. This presents a unique problem and challenge for the fabrication of materials to be used as engineering tissue scaffolds. We believe that electrospinning represents an attractive scaffold processing method to meet the requirements noted. In this chapter, we first discuss how electrospinning can be done in the laboratory and follow with a brief historical summary of the technique. Attention is then directed to work in our laboratories on the use of electrospinning to create a wide variety of scaffold materials for tissue engineering. II THE ELECTROSPINNING SET-UP A typical apparatus (Fig. 1) comprises a short small-bore conduit through which polymer solution is introduced by a suitable pump or from a pressurized reservoir. For reasons that will emerge, the liquid must have some electrical conductivity, i.e., contain anions and cations. Because it is convenient to use stainless steel tubing from which hypodermic needles are produced, the injection conduit is frequently referred to as a “needle.” Opposite the needle at some distance away is a counterelectrode that may have various configurations and compositions ranging from a simple, static metal plate to a rotating vane, drum, or disk. Thus, a potential applied to the injection needle (or to the entering liquid by means of an electrode with which it is in contact) produces an electric field at the needle tip, the intensity of which depends upon the potential difference between the needle and the counterelectrode, the diameter of the needle, and the distance to the counterelectrode. As will be described later, the field induces flow of the emerging liquid toward the counter electrode in the form of a thin jet of liquid. When the applied voltage is sufficient to induce enough charge repulsion to overcome surface tension, a steady stream of fluid is ejected from the tip of needle or pipette. Sufficiently high polymer concentrations in solution favor the creation of a contiguous fiber. In actuality, the technique is even more versatile, as we find that electrospinning can also be done on a dielectric material interposed between the ground and the spinning solution. Thus, a wide variety of substrates other than metals can be coated.

Figure 1 A schematic of an electrospinning system for the production of micro- to nanoscale fibers and subsequent tissue engineering scaffolds.

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During the jet’s travel, the solvent gradually evaporates, and a charged polymer fiber is left that accumulates on the grounded target. It was initially believed that very thin fibers result from splaying of the main fiber jet, although recent studies (discussed later) reveal that a single fiber jet, progressively thinned as the result of whiplike motions, is often deposited on the target. The charges on the fibers eventually dissipate, being neutralized by the surrounding environment. The final product of the process is a nonwoven fiber mat that is composed of tiny fibers with diameters on the order of nanometers to microns [8]. We also note that polymer melts can also be electrospun [9].

III BRIEF HISTORY OF ELECTROSPINNING A Early Work and a General Mechanism The first coherent investigation of the flow of a conducting liquid through a tube maintained at a high voltage with respect to a counterelectrode was initiated by Zeleny [10]. Most of his experiments were with aqueous solutions of electrolytes having relatively high values of electrical conductivity and surface tension but low values of viscosity. Under these conditions the emerging liquid formed a fine spray of highly charged droplets that evaporated rapidly in the surrounding gas (laboratory air in those early experiments). Zeleny was able to observe that as a droplet evaporated it would suddenly break up into a plurality of smaller droplets. This phenomenon had been predicted and characterized in a paper by Lord Rayleigh in 1882 [11]. Rayleigh reasoned that as evaporation reduced the size of a droplet, the charge density increase on its surface due to Coulomb repulsion would overcome the surface tension that held the droplet together. This phenomenon that brings about the disruption of the droplet into smaller droplets is now known as a Rayleigh instability. Many years later, Taylor [12] analyzed what should happen to a droplet of nonconducting, polarizable liquid at the tip of a tube that is subjected to an intense electric field. He found that under no-flow conditions, the competition between electrical forces and surface tension should result in the formation of a liquid cone at the tip of the tube whose base is the diameter of the tube and whose apex angle has the same value for the conical angle of about 50° for all liquids. However, countless experiments show that when the liquid has a finite conductivity, even very small, the stable configuration becomes a cone with a thin column or jet of liquid emerging from the tip. Even so, such “cone-jet” configurations are commonly referred to as Taylor cones, though the conditions under which they are formed depart widely from the assumptions underlying Taylor’s analysis. Except for an occasional paper, the “electrospray” phenomenon described in Zeleny’s experiments remained a laboratory curiosity for many years. In 1968, Dole proposed that charged droplets might provide a way of producing gas phase ions of polymers that could not be ionized by the then available methods, all of which depended upon collisions of gaseous molecules with energetic electrons, photons, or other ions [13]. Dole reasoned that electrospraying a dilute solution of polymer molecules in a volatile solvent could produce charged droplets that would undergo a Rayleigh instability. Each of the offspring droplets would repeat the vaporization–instability sequence. Ultimately, if the original solution was sufficiently dilute, these vaporization–instability sequences would produce droplets so small that each one would contain only a single polymer molecule that would become a gaseous ion by retaining some of its droplet’s charge as the last of the solvent evaporated. For a variety of reasons, Dole’s experiments did not provide enough evidence

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for his hypothesis to persuade other investigators to repeat his experiments. In 1984, Yamashita and Fenn [14] showed that Dole’s procedure would indeed produce ions from small solute molecules that could not be vaporized without decomposition. In 1987 the Yale group showed that what has now become known as electrospray ionization (ESI) can produce intact ions of proteins and other biopolymers with molecular weights up to at least 50,000. It has since been shown that intact ions can be produced from molecules with molecular weights of 100,000,000! Moreover, each of these ions is multiply charged to the point that the mass/charge ratios seldom if ever exceed 3000 or so. Consequently they can be “weighed” with relatively modest mass analyzers. It is now generally agreed that the formation of the charged droplets that are the basis for ESIMS takes place in accordance with the following general mechanism [15]. Fluid emerges from the tip of a small-bore tube at a high potential with respect to an opposing counterelectrode. The liquid forms a so-called Taylor cone from whose tip emerges a thin filament or jet of liquid. For convenience we assume that the needle through which the liquid flows is at a positive potential relative to the counterelectrode. The high field at the tip of the needle drives anions to the needle surface where they give up their negative charge leaving an excess of cations on and near the surface of the liquid jet. The concentration of these excess cations is highest at the surface of the liquid and decreases toward the axis of the jet. Meanwhile the intensity of the external field also decreases exponentially from its maximum at the liquid surface in a way that is characterized by the so-called Debye length, the distance over which the field decreases to 1/e of its value at the surface. The excess cations on and in the liquid are driven by the field toward the counterelectrode, dragging liquid with them. The net result of the decrease in both field strength and concentration of excess cations with distance from the surface is that the outer layers of liquid move toward the counterelectrode with a velocity that decreases rapidly with distance below the surface. Effectively, a thin “skin” of liquid is pulled off the surface of the cone to form the jet of liquid that is accelerated toward the electrode. Meanwhile, the interaction of viscosity and surface tension produces so-called varicose waves on the surface of the jet that increase in amplitude until they truncate the jet into a sequence of uniform charged droplets, another type of Rayleigh instability that is frequently manifested in unconfined liquid columns [16,17] and seen in elongated droplets of polymers in electric fields [18]. Coulomb repulsion disperses the trajectories of the charged droplets into the conical array that forms the characteristic electrospray. Evaporation of solvent from the droplets results in the formation of gaseous ions from the solute species. Electrospinning is effectively an extension of electrospraying and is of particular interest because of the ability to generate polymer fibers of submicron dimensions, down to about 0.05 m [8], a size range that has been heretofore difficult to access yet one which is of great interest for tissue engineering. In electrospinning, polymer solutions or melts are deposited as fibrous mats rather than droplets, with advantage taken of chain entanglements in melts or solutions at sufficiently high polymer concentrations to produce continuous fibers. The electrospinning phenomenon is mechanistically similar to electrospraying, a key difference being that high viscosity and extensive chain entanglements help to stabilize the initial jet toward break-up into droplets, although this does not preclude fragmentation into droplets via a Rayleigh instability. Solvent evaporation (or crystallization from the melt) ultimately stabilizes the jet, the result being a continuous fiber. Early pioneering work was done by Formhals who was granted the first U.S. patent for an electrospinning process that produced fine fibers from a cellulose acetate solution [19]. Baumgarten electrospun acrylics and created fibers of diameters in the range of

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50–1100 nm [20], while Larrondo and Manley described the electrospinning of polypropylene and polyethylene [9,21,22]. B Recent Research In the last decade, there has been increasing interest in studying the details of the electrospinning process and its fibrous products. Our review will not be exhaustive but will attempt to highlight several important developments. For example, Srinivasan and Reneker dissolved poly(p-phenylene terephthalamide) (Kevlar®) in sulfuric acid and were able to obtain fibers with diameters of about 1 m that exhibited crystallization properties similar to that of commercial Kevlar fibers spun from the liquid crystalline state [23]. For these experiments, the fibers were collected in a grounded water bath, which allowed for removal of residual acid. Carbon nanofibers with diameters in the range of 100–500 nm have also been electrospun from polyacrylonitrile solution and molten mesophase pitch [24]. Also, Doshi and Reneker electrospun poly(ethylene oxide) (PEO) as a solution in water and obtained fiber diameters in the range of 0.5–5.0 m [25]. Jaeger conducted atomic force microscopy (AFM) studies on PEO fibers and observed a high level of order in the fiber, with evidence of parallel packing of polymer chains [26]. The voltage and resulting electric field applied to the system influence fiber morphology and diameter, while the orientation of the electric field lines have some effect on final fiber orientation. Since less force and more resistance are applied to draw the solution into a jet, lower voltages lead to thinner fibers. In the same paper [26], a two-electrode electrospinning set-up was employed in which a second electrode is placed concentrically between the nozzle and collecting plate near the nozzle. This was an effort to stabilize the orientation of the electric field and avoid the possibility of corona discharge. Gibson et al. have studied transport properties of electrospun fibers and mats. The nonwoven fiber mats inherently possess very small pores, on the order of microns, yet are still highly porous due to a high density of these tiny pores. Because of these properties, the mats are more resistant to convective gas flow than typical clothing materials such as GoreTex® and cotton but still allow for water vapor diffusion [27,28]. This translates to superior wind resistance and breathability, characteristics that may prove to be very useful for aerosol chemical warfare protective clothing. Since this behavior is more reminiscent of a membrane than a fabric, these properties also suggest the possible advantage of using these mats for filtration. On the mechanistic front, recent studies [29] have indicated that the initial fiber may not splay into thinner ones, but rather progressively thin itself prior to deposition. Thus, electrospinning is essentially the continuous deposition of a single fiber. The experiments used high-speed photography (2000 frames per second) to capture images of the traveling jet. Shortly after a smooth segment was ejected from the nozzle, the jet becomes unstable and bends back and forth until the perturbations become so large that the jet begins to follow a spiraling path in three dimensions. This spiraling path grows in circumference and the jet diameter decreases continually along the path to the collecting plate. At some points along the path, sections of the jet develop a separate and smaller bending instability that follows the same sequence of events, contributing to the decrease in diameter of the main jet. Using the data acquired, a theoretical model that includes the rheological behavior of the solution was developed. Deitzel et al. [30,31] have detailed the effects of solution concentration and spinning voltage on fiber morphology. The diameter, morphology, and orientation of the resultant

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fibers are a function of the many variables involved. The properties of the solution, ranging from the nature and chemistry of the polymer and solvent to the polymer molecular weight and concentration to the solution viscosity and surface tension, are the most changeable parameters. An increase in solution viscosity or concentration gives rise to a larger and more uniform fiber diameter [26,30,31]. Additionally, higher concentrations result in increased surface tension and a more regular, cylindrical fiber morphology. Interestingly, for the case of PEO/water, a bimodal distribution of fiber diameters is observed at high (8 wt%) polymer concentrations. If fibers are still wet when they contact the collecting plate, they will agglomerate to some degree with other fibers in their vicinity, forming “beads” at these junctions. As voltage increases, the number of bead defects increases. This is because the rate at which the solution is removed from the tip of the nozzle exceeds the rate at which the solution is delivered to the tip, and the jet becomes unstable as it is drawn from the inside surface of the nozzle, as opposed to the Taylor cone. Higher polymer concentrations mean less solvent in a given volume of solution, and this leads to drier fibers [30,31]. Solution infusion rate into the system and the distance between the initiation tip and collecting plate also affect fiber dryness. Solvent volatility is another major factor in fiber dryness. In addition to solvent volatility, temperature and makeup of the spinning atmosphere, e.g., relative humidity can affect fiber dryness and, therefore, formation of bead defects. Also, splaying of the main jet was observed under certain conditions [31]. C Electrospinning of Biomaterials and Bioderived Polymers As early as 1977, Martin and Cockshott [32] reported on the use of electrospinning for biomaterials applications with the production of fibrillar mats for wound dressings and vascular prosthetics. They noted that polymer concentration required for spinning depended upon the molecular weight of the polymer, with lower molecular weights requiring higher concentrations. In 1978, Annis et al. [33] described the use of electrospinning to prepare an elastomeric, polyurethane-based vascular prosthesis, and since that time numerous patents have focused on vascular graft materials, with an emphasis on polyurethanes and silicones (see e.g., Refs. 34–38). Martin et al. [39] used electrospinning to prepare porous protein coatings, and Huang et al. [40] electrospun synthetic analogs of elastin. Bognitzki et al. [41] utilized poly(lactic acid) in solution with tetraethyl benzyl ammonium chloride (TEBAC) to electrospin fibers with average fiber diameters of about 1 m. The TEBAC, an organic soluble salt, helped to decrease fiber diameter by increasing solution surface tension and electrical conductivity. Zarkoob et al. reported on the successful electrospinning of silk fibers [42]. Scardino and Balonis [43] discussed the electrospinning of PLA-PGA copolymers and suggest applications in tissue engineering. Efforts in our laboratories have been directed toward the development of scaffolds for tissue engineering using poly(glycolic acid) (PGA), poly(lactic acid) (PLA), polycaprolactone (PCL), collagen, and elastin [44–48], which is the basis of the work summarized in the following section. IV ELECTROSPINNING OF TISSUE ENGINEERING SCAFFOLDS A Poly(glycolic Acid) Poly(glycolic acid) is a biocompatible and bioresorbable polymer that has been a very popular scaffold material since the early days of tissue engineering. The degradation rate of PGA, typically a few weeks, depends on the polymer molecular weight and environmental conditions during the degradation process [49]. It can be processed through traditional ex-

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Figure 2 Electrospun PGA illustrating the fiber diameter (1.2  0.4 m) produced and the overall structure (random fiber orientation) of the scaffold produced (1600, with the inserted size bar at 10 m). trusion techniques (hot melt or cold draw) for products such as sutures or through electrostatic processing to produce nonwoven sheets or tubes. The traditional processing techniques have limited fiber diameters that can be produced, which are in the range of 10–12 m. Electrostatic processing can consistently produce fiber diameters at or below 1 m [48]. By controlling the pick-up of these fibers, the orientation and mechanical properties can be tailored to an individual’s need. A 14 wt% by volume solution was processed to produce 1.2  0.4 m fibers both in an aligned mat and a random mat. Material testing was performed yielding elastic moduli of 105 MPa along the principal fiber axis, 56 MPa orthogonal to the principal fiber axis, and 95 MPa for the random mat. Strain-to-failure was also measured at 64, 106, and 83%, respectively, for the aligned, orthogonal, and random samples. (See Figs. 2 and 3.) B Poly(lactic Acid) Poly(lactic acid) (PLA) is a common tissue engineering scaffold. A common form of PLA is the homopolymer of the L isomer and this is the material used in our laboratories. It has a degradation time of 30–50 weeks depending on the polymer molecular weight and the environment encountered in vitro or in vivo. Like PGA, it can be processed using traditional fiber production methods as have been used in the biodegradable suture market for years. We have determined that the electrospinning of PLA from chloroform at 1/7 w/v (g/mL) is optimal for fibrous mat production [46]. With a static mandrel, the resulting fibrous mat is composed of randomly laid PLA fibers, as illustrated in Fig. 4. The average fiber diameter produced is 10.3  1.3 m. Using a mandrel rotating at approximately 500 rpm leads to the orientation of the fibers along a common axis (Fig. 5). Mechanical testing of PLA electrospun mats indicates a modulus on the order of 180 MPa. PLA is a semicrystalline polymer, which affords this relatively high modulus, but also leads to a fairly rigid, noncompliant tissue engineering scaffold. We have demonstrated that electrospun PLA can be used as a scaffold for human aortic smooth muscle cells [44] and as a key component, with collagen reinforcement, of a novel vascular graft system [46].

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Figure 3 Electrospun PGA illustrating the overall general structure (aligned 1.2  0.4 m fibers) of the scaffold produced (1100, with the inserted size bar at 10 m).

C Polylactic Acid/Polycaprolactone Blends Poly(caprolactone) (PCL) is well known to have a much higher degree of elasticity than PLA. Thus, in order to add elasticity to a PLA-based tissue engineering scaffold, we have blended PCL with PLA, the latter of which we have already shown to have excellent fiber forming characteristics during electrospinning. Using this PLA/PCL blend allows one to create more compliant fibrous structures for tissue engineering. An example of the results

Figure 4 Electrospun PLA illustrating the fiber diameter produced (10.3  1.3 m) and the overall random structure of the scaffold produced on a static mandrel (400, with the inserted size bar at 10 m).

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Figure 5 Electrospun PLA illustrating the fiber diameter produced (10.3  1.3 m) and the overall aligned structure of the scaffold produced on a mandrel rotating at approximately 500 rpm (250, with the inserted size bar at 100 m).

of the electrospinning of PLA/PCL is shown in Fig. 6. For 33.5% PCL blended with PLA (dissolved in chloroform), the average fiber diameter produced is 7.9  1.5 m. At an 8% PCL in PLA blend, the average fiber diameter is 10.3  1.5 m. The mechanical properties of the 33.5% PCL in PLA blend scaffold showed a modulus on the order of 35 MPa and 124 MPa for the 8% PCL blend with PLA. Poly(caprolactone) has a degradation rate that can take 1–2 years, again depending on the polymer molecular weight and environmental conditions during degradation.

Figure 6 Electrospun 8% PCL with PLA illustrating the fiber diameter produced (10.3  1.5 m) and the overall aligned structure of the scaffold produced on a mandrel rotating at approximately 500 rpm (250, with the inserted size bar at 100 m).

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D Collagen Collagen is a major structural protein (native scaffold) present in vertebrates and accounts for approximately 25% of the total protein [50]. For this reason, it is also thought by many in the tissue engineering field to be the “ideal” scaffold material. To date, more than 19 different types of collagen have been described [51]. The main function of collagen is to provide structural support to the tissue in which it is present, but it is also known to sequester many factors required for tissue maintenance and regeneration. The first key issue for electrospinning of collagen was the identification of a solvent that suitably dissolved collagen at sufficient concentrations to accomplish electrospinning. It was also required that the solvent be relatively volatile to ensure rapid drying of electrospun mats. The solvent 1,1,1,3,3,3-hexaflouro-2-propanol (HFIPA) represents a good choice as it is volatile (bp 61°C), and it has been used as a solvent in which protein/amino acid sequences have been suspended for various conformational analysis studies [52,53] and as a solvent for the electrospinning of silk [42]. HFIPA has been shown to induce an -helix conformation of a polypeptide in solution only if the sequence possesses a natural tendency to form the -helix [53]. Collagen electrospinning [47] was performed utilizing type I collagen (calf skin) dissolved in HFIPA (0.083 g/mL). After multiple investigations, the optimally electrospun type I collagen mat under scanning electron microscopy (SEM) revealed a scaffold composed of polymerized collagen fibers with an average diameter of 100  40 nm (Fig. 7). The transmission electron microscopy (TEM) evaluation showed the 100-nm collagen type I fibers with the typical banded appearance characteristic of the native polymerized collagen. It should be duly noted that these small diameter fibers approach the limits of collagen type I fiber formation since the fiber dimension approaches the 50-nm minimal collagen type I fiber (fibril) dimension. It should also be noted that the electrospun mats (Fig. 8) of collagen produced possess substantial structural integrity, which allows them to be removed with care from the mandrel and handled. As an example of the collagen scaffold structural integrity, a collagen mat with an average width of 4.20 mm (longitudinal fiber orientation) and a thickness of 187 m could withstand an average peak load of

Figure 7 Electrospun type I collagen (calf skin) illustrating the fiber diameter produced (100  40 nm) (4300, with the inserted size bar at 1.0 m).

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Figure 8 Electrospun type I collagen (calf skin) illustrating the cross-section of a scaffold produced having a overall thickness of approximately 50 m (850, with the inserted size bar at 10 m). 1.17  0.34 N with a peak stress of 1.5  0.2 MPa. The average modulus for these samples was 52.3  5.2 MPa. Rotation of a collection mandrel during electrospinning affords oriented collagen fibers (Fig. 9). Importantly, transmission electron microscopy (TEM) evaluation shows a continuous 67 nm banding, indicative of natural collagen supermolecular structure [47]. E Elastin The elastic properties of many tissues are due to the presence of elastin in the extracellular matrix and in many tissues the elastin alternates among collagen fibers. We find that pure

Figure 9 Electrospun type I collagen (calf skin) illustrating the fiber orientation from a mandrel that was rotating at approximately 4500 rpm during the electrospinning process to create a scaffold of aligned collagen fibers (430, with the inserted size bar at 10 m).

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Figure 10 Electrospun elastin (bovine ligamentum nuchae) illustrating the 1.1  0.7 m fibers produced (4300, with the inserted size bar at 1 m).

elastin (bovine ligamentum nuchae) can be electrospun at both 20 and 30% (w/v) from HFIPA. At 30%w/v, the elastin displayed optimal electrospinning (fiber formation) properties. Analysis under SEM revealed an average fiber diameter of 1.1  0.7 m (Fig. 10). We have also demonstrated that elastin/collagen blends can be electrospun. V CONCLUSIONS In conclusion, we believe that it may now be possible to construct “biomimicking” fibrous scaffoldings for tissue engineering using the process of electrospinning. With the electrospinning process, scaffolds of various shapes and sizes can be constructed while at the same time precisely controlling fiber orientation, composition (blended fibers), and dimensions. This flexibility in processing is of particular interest when one is trying to mimic the native architecture (multiple layers and orientation of natural fibers) during tissue engineering scaffold development. Another key advantage is that almost all the scaffolds one can make a mold for will be seamless, three-dimensional scaffolds. This elimination of seams will eliminate any variation in the scaffold or possible weak areas during tissue development/ regeneration. All the scaffolds also exhibit substantial structural strength for ease of handling. Also, the ability to co-spin polymers with various additives (e.g., growth factors, DNA) to control their slow release [54] offers the possibility of tuning the function of the scaffold toward selected cell types. Thus, with the broad ability to electrospin both natural and synthetic polymers, it may now be possible to create ideal tissue engineering scaffolds that are customized for specific tissue growth. Current work in our laboratories is focusing on the creation of a broad range of tissues in vitro and in vivo using electrospun scaffold platforms. ACKNOWLEDGMENTS The authors would like to thank the Whitaker Foundation (RG-98-0465) and NASA Langley for the support of this research as well as Alkermes, Inc., for the donation of PGA, PLA,

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and PCL. We would also like to thank Ms. Judy Williamson for obtaining many of the SEM micrographs. REFERENCES 1. How T. V., Guidoin R., Young S. K. 1992. Proc. Inst. Mech. Eng. 206:61. 2. Vacanti J. P., Vacanti C. A. 1997. In: Principles of Tissue Engineering, Lanza R., Langer R., Chick W., Eds. R. G. Landes Co. 3. Greisler H. P. 1996. Biomaterials 17:3. 4. Mooney D. J., Langer R. S. 1995. The Biomaterials Handbook, CRC Press. 5. Starke G. R., Douglas A. S., Conway D. J. 1999. In: Tissue Engineering of Prosthetic Vascular Grafts, Zilla P., Greisler H. P., Eds. R. G. Landes Co. 6. Hsu S., Kambic H. 1997. Artif. Organs 21:12. 7. Olsen B. R. 1997. In: Principles of Tissue Engineering, Lanza R., Langer R., Chick W., Eds. R. G. Landes Co. 8. Reneker D. H., Chun I. 1996. Nanotechnology 7:216. 9. Larrondo L., St. John Manley R. J. 1981. Polymer. Sci. Polym. Phys. Ed. 19:909. 10a. Zeleny J. 1914. Phys. Rev. 3:69–91. 10b. Zeleny J. 1917. Phys. Rev. 10:1–6. 11. Rayleigh J. W. G. 1882. Lond. Edinburgh Dublin Phil. Mag. 14:184. 12. Taylor G. I. 1969. Proc. Royal Soc. Lond. Ser. A 313:453. 13. Dole M., Hines R. L., Mack L. L., Mobley R. C., Ferguson L. D., Alice M. B. 1968. Macromolecules 1:96. 14a. Yamashita M., Fenn J. B. 1984. J. Phys. Chem. 88:4451–4456. 14b. Yamashita M., Fenn J. B. 1984. J. Phys. Chem. 88:4671–4675. 15. Maekawa M., Nohmi T., Zhan D., Kiselev P., Fenn J. B. 1999. J. Mass Spectrom. Soc. Jpn. 47:76. 16. Rayleigh J. W. G. 1945. The Theory of Sound, Vol. 2, Dover; New York. 17. Adamson A. W., Gast A. P. 1997. In: Physical Chemistry of Surfaces, 6th ed., Wiley-Interscience; New York, pp. 9–10. 18. Serpico J. M., Wnek G. E., Krause S., Smith T. W., Luca D. J., Van Laeken A. 1991. Macromolecules 24:6879. 19. Formhals A. U.S. Patent 1,975,504. (1934). 20. Baumgarten P. K. 1971. J. Coll. Interface Sci. 36:71. 21. Larrondo L., St. John Manley R. 1981. J. Polym. Sci. Polym. Phys. Ed. 19:921. 22. Larrondo L., St. John Manley R. 1981. J. Polym. Sci. Polym. Phys. Ed. 19:933. 23. Srinivasan G., Reneker D. H. 1995. Polym. Int. 36:195. 24. Chun I., Reneker D. H., Fong H., Fang X., Deitzel J., Tan N. B., Kearns K. 1999. J. Adv. Mater. 31:36. 25. Doshi J., Reneker D. H. 1995. J. Electrostatics 35:151. 26. Jaeger R., Bergshoef M. M., Martin C., Batlle, Vancso G. J. 1998. Macromolecular Symposia 127:141–150. 27. Gibson P., Schreuder-Gibson H., Pentheny C. 1998. J. Coated Fabrics 28:63. 28. Gibson P. W., Schreuder-Gibson H., Rivin D. 1999. AIChE J., 45:190. 29. Reneker D. H., Yarin A. L., Fong H., Koombhongs S. 2000. J. Appl. Phys. 87:4531. 30. Deitzel J. M., Kleinmeyer J., Harris D., Beck Tan N. C. 2001. Polymer 42:261. 31. Deitzel J. M., Beck Tan N. C., Kleinmeyer J. D., Rehrmann J., Tevaul D. 1999. Army Research Laboratory Technical Report, ARL-TR-1989. 32. Martin G. E., Cockshott I. D. 1977. U.S. patent 4,043,331. 33. Annis D., Bornat A., Edwards R. O., Higham A., Loveday B., Wilson J. 1978. Trans. ASAIO 24:206. 34. How T. V. 1985. U.S. patent 4,552,707.

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10 Clinical and Biomechanical Design Considerations for Engineered Cortical Bone Allografts Kai-Uwe Lewandrowski, Francis J. Hornicek, Frank X. Pedlow, Mark C. Gebhardt, Henry J. Mankin, and William W. Tomford Massachusetts General Hospital, Boston, Massachusetts

I INTRODUCTION Allograft tissue transplantation has become common practice in orthopedic surgery. Procedures range from soft tissue allograft for reconstruction of knee ligaments to allograft bone struts and corticocancellous chips to re-establishing bone stock in total joint reconstruction. Structural allograft reconstruction after bone tumor resection has met with varying success [1,2–4]. These types of bone grafts, however, have been found to incorporate slowly into host bone resulting in susceptibility to nonunion, fatigue fracture, and infection [1,5–10]. As a result, complications associated with the use of structural allografts are not infrequent and have caused some surgeons to abandon their use in favor of large metal prostheses for skeletal reconstruction after tumor resection [11,12]. These prostheses have a lower initial complication rate, but tend to fail with time [6–9]. Structural allografts have certain advantages over metal endoprostheses that are well recognized [13]. Investigators from Massachusetts General Hospital [3,4] previously reported on the outcome and complications of osseous reconstructions using large allografts. The earlier report indicated a fracture rate of 10.4%, whereas the later report identified a fracture rate of 16%. The study of the fracture pattern with large cortical allografts has served as the basis for this review of clinical and biomechanical design considerations. These patterns have to be taken into account when developing bone allograft with improved incorporation into host bone. Commonly used methods for improving osteoconductive and osteoinductive properties of bone allografts, including demineralization and increasing porosity, could 179

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have detrimental effects on clinical outcomes due to mechanical failure of the graft if the complexity of multiple factors affecting the allograft transplantation process are not addressed. It is the authors’ purpose to review how methods of enhancing osteoconduction and osteoinduction can be applied to cortical bone allografts while assuring success of the clinical transplantation procedure. II CURRENT METHODOLOGY Patients treated for musculoskeletal tumors at Massachusetts General Hospital for bone resections and structural allograft reconstruction resulted in a total of 1046 allograft implantations between 1974 and 1998. The allografts were obtained from the bone bank at the hospital. All allografts were obtained, processed, stored, and evaluated for bacterial infection and other blood-borne transmissible disease according to the requirements of the American Association of Tissue Banks [10,14–16]. All patients received intravenous antibiotics preoperatively and postoperatively until discharge. They then were maintained on oral antibiotics for 3 months. If the resection involved the lower extremity, warfarin was given for 6 weeks. Patients having lower extremity reconstruction were placed in a long leg cast prior to their discharge from the hospital. The patients were kept partial weightbearing. After the soft tissues had healed, the extremity was placed in a knee–ankle–foot orthosis and the patient was allowed to start ankle and knee range of motion (ROM) exercises. The patients were not allowed to bear full weight until after graft–host junction had healed. Patients who had resections around the hip or humerus wore a brace to immobilize that part. At 6 weeks postoperative, patients who had surgery around the upper extremity started ROM exercises without wearing the brace. Serial radiographs were obtained of all patients to assess healing. Union is defined as new bone formation across the host–donor junction and the absence of a radiolucent line. A nonunion is arbitrarily defined as the failure of the host–donor junction to unite 1 year after the index procedure. Any hardware failures are considered to be caused by a nonunion, not a fracture. In this clinical review, a functional score as reported by Mankin et al. [17] was assigned after follow-up of the allograft fracture. The rating scale was as follows: Excellent—a patient with no evidence of disease, no pain, and essentially normal function with no limitations other than high performance activities. Good—a patient with no evidence of disease, but with some degree of impairment of function that did not necessitate the use of a brace or ambulatory aids or interfere with the patient’s occupation or lifestyle. Fair—patients with no evidence of disease, but who required a brace or ambulatory aid and had sufficient pain that impaired function. Failed—patients who died, required amputation, or in whom the graft was removed.

III MECHANICAL FAILURE OF BONE ALLOGRAFTS Three major types of allograft fractures were identified. The incidence of nonunion of the allograft and host–donor bone junction was found to be increased in patients with fractured allografts. More current reports [18,19] on the complication of allograft fracture in a larger series of allografts transplanted for large structural defects made by tumor resection indicated an increase in the fracture rate in patients who received chemotherapy.

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However, no correlation was found between chemotherapy or radiation therapy and allograft fracture. Irradiation of the graft before implantation predisposes to fracture [20]. In addition, the authors were unable to identify any difference between skeletal site and allograft fracture rate, but observed a difference between the type of allograft and rate of fracture. The osteoarticular allografts and allograft arthrodeses were more likely to fracture. The allografts used for arthrodesis have higher stresses than those used for most other applications. Examining the intercalary and composite grafts reveals that the entirety of the allograft in these reconstructions is supported by hardware or prostheses. This is in contrast to the osteoarticular grafts in which the articular end of the graft often is unsupported for some distance. Allograft fractures have been classified into three types by Berrey et al. [1]. Type I fractures result in rapid dissolution of the allograft, probably are immunogenically related, and are uncommon. Type II fractures involve fracture of the shaft of the allograft, occur at a mean of approximately 28 months postoperatively, and are more probable when the allograft is fixed with plates and screws, rather than by intramedullary fixation. Type II fractures sometimes heal with nonoperative treatment. In the series of Berrey et al. [1] the fractures in two of 22 patients healed with nonoperative treatment. A new allograft was placed in three patients; in one an arthrodesis of the wrist was done with autogenous iliac crest bone graft after removal of the allograft; and in 16 patients open reduction and internal fixation of the fracture was performed. The fractures in 6 of these 16 patients healed without additional intervention, whereas the other 10 patients required multiple additional procedures to regain function. Type III fractures consist of fragmentation of the joint, occur later (mean of 32 months postoperatively), and often are salvaged with total joint replacement forming an allograft prosthetic reconstruction, a revision allograft procedure, or internal fixation and bone grafting. Vander Griend [21] reported that the type and quality of internal fixation was related to the risk of fracture and allograft nonunion. Twelve fractures occurred in his series of 120 allografts. All of the fractures were type II and occurred in relation to screw holes or a bony defect in the allograft. Quality of the allograft–host junction was associated with the healing rate. Gaps greater than 2 mm resulted in slower healing or a higher incidence of nonunion. Stable fixation at the graft–host junction also facilitated union. IV CLINICAL SIGNIFICANCE Of the 1046 patients reviewed for the purpose of this study, 185 had fractured their allograft. Patient demographics, site of reconstruction, stage of disease, allograft type, whether adjuvant therapy was used, whether the allograft was radiated for positive cultures after procurement, graft length, type of graft fixation, whether the fracture involved a screw hole, other associated complications, the treatment of the fracture, and ultimate outcome have been analyzed. There were 183 patients whose structural allografts fractured. In two of these patients, new allografts that had been implanted to treat the first fracture subsequently fractured. This yielded 185 fractures in 1046 grafts for an incidence of 17.7%. The average age of the patients was 29.8 years. There were 97 females and 86 males. The mean time to fracture was 3.2 years. Fifty-three patients had benign tumors, 11 had stage IA disease, 28 had stage IB disease, 71 had stage IIB disease, and 20 had stage III disease. There was no significant difference between the patients whose allografts fractured and patients whose allografts did not fracture regarding gender, age, or stage of disease.

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The initial fixation of the allografts included plates and screws in 181 patients, intramedullary nails in two, and combination intramedullary nail and plate in two. There were eight type I fractures, 114 type II fractures, and 63 type III fractures. Sixty-one of the fractures involved a screw hole (33%). A typical type II fracture that occurred at the end of the metal plate and through a screw hole is shown in Fig. 1. A different kind of type II fracture is shown in Fig. 2. The fracture line extends from the graft–host junction distally in line with the screws. Data on the length of the fractured allografts were available for 121 of the allografts. The average length of these grafts was 15.5 cm. Functional outcome was evaluated in relation to length of graft in patients with fractures. Sixty-four grafts were less than 15 cm and 57 were greater than 15 cm. If the length of the allograft was longer than the average, the patient tended to do worse after fracture than if the length was less than the average (p  0.3). Eighty-eight of the patients whose allografts fractured received adjuvant therapy. Sixty-one of the patients had chemotherapy, 15 had radiation therapy, and 12 had both. Ninety-five patients had no adjuvant therapy. When compared with the patients whose allografts did not fracture, adjuvant therapy had no effect on the fracture rate (p  0.05). Deterioration in function for those patients who received chemotherapy and whose allografts fractured compared with those patients who did not have chemotherapy was found. Radiation therapy did not have this effect.

Figure 1 A type 2 allograft fracture is shown in a patient who had a distal femoral reconstruction 4-years previously. The allograft–host bone junction is well healed.

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Figure 2 A type 2 allograft fracture is shown in a patient who underwent allograft reconstruction after osteosarcoma resection. The fracture occurred 4 weeks after bone transplantation.

V MANAGEMENT STRATEGIES Current management includes nonoperative treatment for fractures if they are minor and the displacement is minimal. However, the success rate is low. Allograft-to-allograft healing is extremely rare, but has been recently reported in a 10-year-old boy with a stage IIB [7] osteosarcoma of the diaphysis of the femur. The lesion involved the central half of the femur and a large soft tissue mass extended into the vastus intermedius. The patient’s chest was clear. He received neoadjuvant chemotherapy and underwent resection and reconstruction with an intercalary allograft. The distal graft–host junction failed to unite. The remaining distal femur was removed and a second allograft was inserted to create an arthrodesis of the knee. The arthrodesis was performed using a Sampson nail (3-M, St. Paul, MN) and resulted in an allograft–allograft junction in the femur and an allograft–host a junction in the tibia, both of which healed in 7 months. The patient did well, until he had a fracture develop through the midportion of his second allograft and failure of the intramedullary rod. The reconstructive options considered included revision of the arthrodesis with replacement of the entire allograft (to the lesser trochanter), revision of only the distal fractured allograft, rotationplasty, or amputation. Eventually, the patient underwent revision of the allograft arthrodesis. The Sampson intramedullary nail was removed and replaced with an interlocking intramedullary nail (Richards, Memphis, TN). The distal portion of the allograft was removed and replaced with a new allograft. At the previous allograft–allograft junction a circumferential ring of

Figure 3 (A) Anteroposterior (AP) radiograph of the femur showing the osteoblastic tumor involving the diaphysis of the femur. (B) Anteroposterior view of the femur showing an intercalary allograft with early union at the proximal allograft–host junction and failure at the distal femoral allograft–host junction. (C) Anteroposterior radiograph of the leg of the patient in Case 2, 14 months after the revision arthrodesis procedure. The proximal allograft–allograft junction and the distal allograft–host junction have healed. (D) The AP view shows the revision allograft–arthrodesis with healing at (A) the proximal allograft–host junction, (B) the allograft–allograft junction, and (C) the distal allograft-host junction. The CT scan at Level B shows union of the allograft–allograft junction evident on the AP and axial views. 184

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fracture callous was preserved and the new allograft was inserted into this sleeve of viable bone, forming an allograft–allograft junction surrounded with fracture callous. At 14 months postoperatively, CT scans revealed allograft–allograft union (Fig. 3). Other options to manage fractured allografts include intramedullary fixation; careful allograft–host size matching and tight allograft–host junctions with solid internal fixation have reduced the incidence of allograft fracture and nonunion. Liberal use of rotational muscle flaps has decreased the risk of wound healing complications and the overall infection rate. The technique of Capanna et al. [22] goes one step further by proactively adding a vascularized fibula in conjunction with the allograft to help minimize the risk of allograft fracture and facilitate healing when fractures occur. Studies are ongoing regarding allograft–host immunologic matching to decrease the risk of type I fractures and perhaps even some type II and III fractures. VI APPROACH TO ENHANCING INCORPORATION Several decades of experience with massive allografts have resulted in some improvements in the rates of allograft fracture, nonunion, and infection [13]. Recognizing that complications of the allograft procedure are probably multifactorial in origin and depend on factors such as the type of graft, and the use of adjuvant chemo- and radiation therapy, one approach by which the clinical outcome with the allograft system could be further improved is to enhance their incorporation into the host’s own bone. Studies have attempted to improve host incorporation by altering the geometric surface configuration of cortical bone [23–28]. The mechanisms by which the presence of laser holes promoted osteogenesis and incorporation in partially demineralized grafts include the greater surface area of partially demineralized bone and/or increased access to vascular tissue. Previous studies have indicated that the geometry of an implant may influence the extent of bone formation. The osteoconductive potential of cortical bone allografts so treated has been attributed, at least in part, to its morphometric similarity to cancellous bone. Gendler [26] used fully demineralized diaphyseal allogeneic struts perforated by use of a mechanical drill. In comparison, O’Donnell et al. [27] used demineralized calvarial cortical bone. Bernick et al. [23] characterized the inductive cellular events in a similar system. Scanlon [27] implanted demineralized canine femoral strut allografts in an orthotopic model. The latter two studies have used an erbium: yttrium-scandium-gallium-garnet (Er: YSGG) laser for drilling of cortical bone allografts, thereby increasing the bone porosity and allowing demineralization to proceed in areas normally inaccessible to the demineralization process. When reimplanted, these grafts may therefore be more osteogenic than cortical grafts without holes. An advantage of the Er:YAG laser over mechanical drilling is that thermal bone damage does not occur. Tissue is removed by explosive vaporization [29–30]. Furthermore, laser ablation is a noncontact process permitting sterile handling of the grafts. Finally, it allows the drilling of a large number of very small holes in a uniform pattern. Although demineralization has been successful in improving osteoinductive properties of bone in experimental [31–34] and clinical studies [35], fully demineralized cortical bone has found limited clinical application in long bone reconstruction. The primary reason for this is that cortical bone, once subjected to extensive demineralization, loses its essential biomechanical properties. However, sufficient mechanical strength of these grafts is necessary in order to withstand skeletal forces when used in long bone reconstruction.

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Demineralization of cortical bone yields a geometric surface configuration that is less advantageous for bony ingrowth when compared with cancellous bone because cortical bone is less porous and has a comparatively low surface area–to–volume ratio. These hypotheses have stimulated the investigation of the process of cortical bone demineralization at the author’s institution [36–40]. Mathematical models predicting demineralization kinetics were developed, which in turn allowed control of mechanical properties of such treated grafts by deriving mechanical models. Good fit to mechanical models was determined by measuring flexural rigidity and compression strength of partially demineralized and laser-perforated diaphyseal bone allografts. These methods were employed in orthotopic transplantation studies in rats and sheep and have been tested clinically in patients who underwent skeletal reconstruction after resection of primary bone tumors or metastatic disease. These studies showed improved incorporation of partially demineralized and laser-perforated cortical bone allografts and suggested that both processes may be applicable to clinical bone allografts. However, there are several clinical concerns related to acceleration of graft resorption and subsequent replacement with the recipient’s own bone, all of which are primarily related to premature mechanical weakening of the allograft. VII ENGINEERED BONE ALLOGRAFTS FOR CLINICAL USE The transplantation studies in rats and sheep have demonstrated that laser-perforated and partially demineralized cortical bone allografts showed the most promising results with respect to incorporation into host bone and mechanical strength of the allografted recipient bone. No pathological fractures were observed in either in vitro study suggesting that our mathematical models allow production of grafts strong enough to withstand skeletal forces when used in long bone reconstruction. In comparison to bone allografts in current clinical use, grafts used in our clinical trials were stripped of soft tissues and the bone marrow was removed. Presumably, cellular sources of allostimulation had been eliminated. On the basis of our radiographic, histologic, and biomechanical observations, we have developed laser-perforated and partially demineralized bone allografts for clinical use (Fig. 4).

Figure 4 Photograph of a clinical intercalary cortical bone allograft processed by laser perforation and subsequent partial demineralization.

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All allografts were obtained, processed, stored, and evaluated for bacterial infection and other blood-borne transmissible disease according to the requirements of the American Association of Tissue Banks [10,14–16]. Before transplantation, grafts were stripped of soft tissues and the bone marrow was removed. The grafts were prepared at various lengths ranging from 30 mm for vertebral replacement to 30 cm for intercalary reconstruction. In spinal applications, grafts were used as interbody spacers. In long bone reconstruction, fixation was achieved with plates and screws. An Er:YAG laser (Schwartz Electro-Optics, Concord, MA) operating at a wavelength of 2940 nm was used to drill holes 300 m in diameter through the entire thickness of the diaphyseal cortex producing a hole that was typically 3 to 6 mm deep. The laser was connected optically to an OPMI 1 F/C operating microscope (Zeiss, Oberköchen, Germany) by an articulated arm. A 100- to 200-s pulse train consisting of pulses 1 s long was used for drilling. The energy delivered per pulse was measured using a pyroelectric joulemeter (ED-200; Gentec, Ste. Foy, Quebec) and oscilloscope. A collinear helium–neon laser with a visible beam was used to aim the Er:YAG laser. The beam diameter at the tissue surface was 330 m as defined by the e2 intensity points. The lasers were aimed by means of a micromanipulator (Laser Mechanisms, Southfield, MI). Typically, drilling was performed using a 53-mJ pulse, with 25 to 30 pulses required to drill each hole. The fluence per pulse was typically 60 J/cm2. Laser holes were drilled at a distance of 2.5 mm between one another. Typically, holes were drilled around the circumference of the graft through the entire thickness of the cortex at a spacing of 10 to 15 mm. Thereafter, demineralization was performed. The methods employed for controlled demineralization have been developed by the authors and are based on the theory of the “classic shrinking-core reaction model,” which has been generally applied for modeling purposes of fluid–solid diffusion systems. This model implies the existence of a sharp reaction front that, in the case of demineralization of cortical bone, separates the demineralized from the mineralized portion. The reaction front is assumed to advance with increasing immersion time in the decalcifying agent, which implies the kinetics of this reaction front can be predicted for some ideal geometries by mathematical models. In turn, this concept allowed implementation of the concept of controlled demineralization to bone grafts in current clinical use. In further studies, the authors determined the demineralization kinetics of human cortical bone and cortical bone of various animal species as the advance of the demineralization front versus immersion time by measuring extraction of bone mineral in both planar and cylindrical geometries [38]. For this clinical trial, the authors employed mathematical models based on diffusional mass transfer to controllably predict the extent of demineralization of the ovine cortical bone allografts and, hence, their mechanical properties prior to transplantation [36–40]. For demineralization, the bone ends were capped with an acid resistant paraffin layer to allow demineralization to proceed only from the periosteal surface to the endosteal surface [37]. The outer dimensions of each graft were recorded and the outer diameter was used to calculate for each individual graft the time needed in the hydrochloric acid bath to achieve a demineralization extent to produce a uniform 20% reduction of the rigidity of the original bone [39–40]. The 20% reduction of the rigidity of the original bone was chosen on the basis of the author’s prior mechanical analyses of partially demineralized bones. Bone allografts that were laser drilled before demineralization were previously measured to have an average reduction of flexural rigidity of 40% if demineralization was added as indicated above. Therefore, laser perforated and partially demineralized allografts would be left with 60% of the rigidity of the original bone [36].

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After this processing allografts were transferred to Osteotech, Inc. (Shrewsbury, NJ) for further processing, including culturing and packaging. Grafts with positive cultures were discarded and not used for this study. All grafts were stored at 80°C until use. Prior to transplantation, allografts were thawed in a solution containing polymyxin B sulfate (500,000 U/l) and Bacitracin (50,000 U/l). VIII CLINICAL APPLICATIONS AND PRELIMINARY RESULTS Cortical bone allografts processed by laser perforation and demineralization were used for spinal and long bone reconstruction. For this study 26 patients were followed retrospectively for a minimum of 2 years (range 24 to 36 months). These patients had undergone anterior spinal column reconstruction after resection of metastatic disease (14 patients; Fig. 5), or primary bone lesions (12 patients) using structural allografts in conjunction with anterior fixation alone or with anterior and posterior fixation (Figs. 6 and 7). Patients with spinal allograft reconstructions were followed clinically and their spinal fusion was assessed using a radiographic grading system: Grade I—fused with remodeling and trabeculae Grade II—graft intact, not fully remodeled and incorporated; no lucencies Grade III—graft intact, but a definite lucency at the top or bottom of the graft Grade IV—resorption of bone graft with or without collapse Of the 26 patients, only three patients underwent revision for local tumor recurrence. Another patient had revision of anterior fixation because of hardware failure. Most patients (21 patients) showed successful spinal fusion with remodeling and incorporation of the graft in the absence of lucencies. Only one patient demonstrated radiographic evidence of nonunion.

Figure 5 On the left, the posterior exposure of the L2 vertebral segment of a patient with metastatic disease is shown after circumferential resection and placement of spinal instrumentation with pedicle screws. On the right, the anterior exposure of the L3 segment of a patient with metastatic disease is shown after anterior resection and vertebral body reconstruction with a laser-perforated and partially demineralized cortical bone allograft prior to placement of instrumentation.

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Figure 6 Anteroposterior and lateral radiograph of the lumbar spine showing a vertebral reconstruction using a 30-mm long laser-perforated and partially demineralized cortical bone allograft with spinal instrumentation. Results of this study show that anterior structural spinal allografts are an effective adjunct when used in the proper clinical situation. No mechanical failures were observed with the allografts used in patients who received them as part of their spinal reconstruction. The extent of increasing porosity of the allografts by perforation and demineralization proved acceptable for these spinal applications. Modified allografts used in this clinical trial showed a clear advantage over metallic interbody spacers with or without bone graft. Perforated and demineralized bone allografts used provided immediate strength, are available in sufficient quantity, and based on the results of this study appear to have fusion rates comparable with autografts, especially if combined with posterior instrumentation. Other applications of perforated and demineralized grafts included transplantation in the reconstruction of long bones either in the form of hemicortical struts (Fig. 8) or complete intercalary allograft sections (Fig. 9). These grafts were processed in such a way that only one-half of the graft was treated by perforation and demineralization. Thus, grafts would serve as their own internal control. The majority of grafts in these patients were anchored with plates and screws. A total of eight patients received a perforated and demineralized cortical bone allograft for their long bone reconstruction. One of the eight allografts showed a nonunion at the host bone–allograft junction site. None of the grafts fractured. A significant decrease in the fracture rate of allografts fixed with intramedullary fixation compared with those fixed with plates and screws has been reported [21]. In another series it was reported that chemotherapy and plate fixation increased the rate of fracture [19]. Al-

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Figure 7 Coronal (top left), saggital (top right), and axial (bottom left) CT Scans, and AP radiograph (bottom right) of the lumbar spine showing a vertebral reconstruction using a 30-mm long laser-perforated and partially demineralized cortical bone allograft with spinal instrumentation in a patient with metastatic disease to L3.

Figure 8 The left photograph shows resection of a portion of the tibia of a 50-year-old male with squamous cell carcinoma of the skin that eroded into the underling bone. Reconstruction was achieved with a 80-mm long laser-perforated and partially demineralized hemicortical allograft strut that was fixed with two screws (right).

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Figure 9 Anteroposterior radiograph of the femur showing reconstruction with an alloarthrodesis using a perforated and demineralized intercalary allograft after resection of an osteoblastic tumor involving the distal femur and the proximal tibia (left). The proximal half of the graft was untreated and the patient had a fracture through the proximal junction site. The distal portion of the allograft was perforated and demineralized and showed radiographic evidence of healing at the tibial junction site with the host bone (second left). The clinical photographs are shown after retrieval of the graft from the patient because of recurrent disease. The proximal untreated junction site (near right) and distal treated junction site (far right) are shown. Please note the fracture of the proximal untreated half of the allograft near the junction site. In comparison, the treated distal half of the allograft had healed with the host bone (far right).

lograft fractures usually occur between 1 and 3 years after implantation and most occur through a screw hole [41]. In analogy, this raises the concern that the laser holes lead to abnormal stress concentration and predisposes the allograft to fracture. However, this mode of failure was not observed in this limited pilot study. IX DISCUSSION As shown by the preliminary data of our clinical trial studies, perforated and demineralized cortical bone allografts can be successfully applied to reconstruction of spine or long bones despite the reduction of mechanical strength that is caused by these two processing techniques. However, it seems clear that a balance between enhancement of incorporation with accelerated graft resorption and loss of the initial and long-term biomechanical properties is needed. Ideally, the modified allograft should provide for rates of bone resorption commensurate with new bone formation during the process of bone allograft incorporation. Clinically applied laser perforation and demineralization results in approximately 40% loss in the overall mechanical properties of the graft at the time of transplantation. The first issue is in addressing the timing and extent of new bone formation. In addressing this, one can invoke the stages of bone induction within the surrounding host bed and then the allo-

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graft in analogy to the description by Bernick [23]: Stage of induction of mesechymal cells of the host bed: Stage of fibrous and vascular tissue invasion: Stage of de novo bone formation in tissues surrounding the graft: Stage of increasing incorporation by bone resorption and new bone formation:

Up to about 1 week To ~2–4 weeks To ~4–6 weeks Several months to years

When bone formation at the junction sites is complete in an average of about 6–12 months, the graft is considered to be clinically and radiographically healed. Once new bone is formed at the graft site, immature bone will start to respond to stresses by remodeling to highly structured bone as indicated by formation of secondary osteons. Depending on the site of transplantation, the quality of the host bed, means of fixation, and other concurrent factors such as the presence of infection, chemotherapy, or radiation, successful incorporation is evident both clinically and radiographically with the majority of the bone graft being incorporated in about 8–12 months. Under ideal conditions, the rate of mechanical losses in the composite should equal the rate of mechanical recovery in the bone throughout the course of bone healing. From the collective experiments [36–40] and preliminary clinical data, it appears feasible to suggest that the initial flexural rigidity reductions of the demineralized/laser-perforated graft did not exceed 40%. Furthermore, using the preceding and assigning values to the initial composite mechanics and to the rate of mechanical loss in the early in vivo course after transplantation, one should be able to assign clinical success criteria on the basis of an overall limitation of mechanical strength to no more than 25–30% at any given time during the incorporation process. The application of perforation and partial demineralization should result in an overall enhancement of bone allograft incorporation for two reasons. First, the outer demineralized surface layer provides a highly osteoinductive matrix for the ingrowth of host mesenchymal cells from the adjacent soft tissues into the allograft that can be transformed in bone forming chondrogenic and osteogenic cells. Second, it may further promote the osteoinductive concept by development of focal centers of new bone formation in patent holes due to the release of growth factors from the demineralized bone matrix into the osteoconductive polymer scaffolds. The cortical bone transplants prepared with these methods are expected to incorporate faster, having the potential to develop increasing porosity and osteoinductive potential over time [24].

ACKNOWLEDGMENTS This work was supported by NIH Grant AR-21896, the Medical Free Electron Laser Program under Office of Naval Research Contract N00014-91-0084, by DOE Grant DEFG02-91 ER61228, and NIH/NIAMS Grant AR-45062 (to Kai-Uwe Lewandrowski). No ownership or profit has been derived from this study.

REFERENCES 1. Allan D. G., Lavoie G. J., McDonald S., Oakeshott R., Gross A. E. 1991. Proximal femoral allografts in revision hip arthroplasty. J. Bone Joint Surg. 73B:235–240.

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2. Musculo D. L., Petracchi L. J., Ayerza M. A., Calabrese M. E. 1992. Massive femoral allografts followed for 22 to 36 years: report of six cases. J. Bone Joint Surg. 74B:887–892. 3. Ottolenghi C. E. 1972. Massive osteo and osteo-articular bone grafts: technique and results of 62 cases. Clin. Orthop. 87:156–164. 4. Parrish F. F. 1973. Allograft replacement of all or part of the end of a long bone following excision of a tumor: report of twenty-one cases. J. Bone Joint Surg. 55A:1–22. 5. Berry B. H., Lord C. F., Gebhardt M. C., Mankin H. J. 1990. Fractures of allografts, frequency, treatment, and end results. J. Bone Joint Surg. 72A:825–833. 6. Bradish C. F., Kemp H. B., Scales J. T., Wilson J. N. 1987. Distal femoral replacement by custom-made prosthesis: clinical follow-up and survivorship analysis. J. Bone Joint Surg. 69B: 276–284. 7. Brien E. W., Terek R. M., Healy J. H., Lane J. M. 1994. Allograft reconstruction after proximal tibial resection for bone tumors: an analysis of function and outcome comparing allograft and prosthetic reconstruction. Clin. Orthop. 303:116–127. 8. Chao E. Y. 1989. A composite fixation principle for modular segmental defect replacement (SDR) prostheses. Orthop. Clin. North Am. 20:439–453. 9. Chao E. Y., Sim F. H. 1985. Modular prosthetic system for segmental bone and joint replacement after tumor resection. Orthopedics 8(5):641–651. 10. Doppelt S. H., Tomford W. W., Lucas A. D., Mankin H. J. 1981. Operational and financial aspects of a hospital bone bank. J. Bone Joint Surg. 63A:1472–1479. 11. Eckardt J. J., Eilber F. R., Rosen G., Mirra J. M., Dorey F. J., Ward W. G., Kabo J. M. 1991. Endoprosthetic replacement for Stage IIB osteosarcoma. Clin Orthop. 270:202–212. 12. Eckardt J. J., Matthews J. G., Eilber F. R. 1991. Endoprosthetic reconstruction of the proximal tibia. Orthop. Clin. North Am. 22:149–159. 13. Hornicek F. J., Gebhardt M. C., Sorger J. I., Mankin H. J. 1999. Tumor reconstruction. Orthop. Clin. North Am. 30:673–684. 14. Friedlaender G. E., Tomford W. W. 1991. Approaches to the retrieval and banking of osteochondral allografts. In: Bone and Cartilage Allografts, Friedlander G. E., Goldberg V. M., Eds. The American Academy of Orthopaedic Surgeons, Park Ridge, IL. 1:185–192. 15. Strong D. M., Sayers M. H., Conrad E. U. 1991. Screening tissue donors for infectious markers. In: Bone and Cartilage Allografts, Friedlaender G. E., Goldberg V. M., Eds. The American Academy of Orthopaedic Surgeons, Park Ridge, IL. 1:193–209. 16. Kagan R. J. 1999. Standards for tissue banking. American Association of Tissue Banks, McLean, VA. 17. Mankin H. J., Doppelt S., Sullivan T. R., Tomford W. W. 1982. Osteoarticular and intercalary allograft transplantation in the management of malignant tumors of bone. Cancer 50:613–630. 18. Hornicek F. J., Mnaymneh W, Lackman R. D., Exner G. U., Malinin T. I. 1998. Limb salvage with osteoarticular allografts after resection of proximal tibia bone tumors. Clin. Orthop. 352: 179–186. 19. Thompson R. C., Pockvance E. A., Garry D. 1993. Fractures in large-segment allografts. J. Bone Joint Surg. 75A:1663–1673. 20. Gebhardt M. C., Roth Y. F., Mankin H. J. 1990. Osteoarticular allograft reconstruction in the proximal part of the humerus after excision of a musculoskeletal tumor. J. Bone Joint Surg. 72A:334–344. 21. Vander Griend R. A. 1994. The effect of internal fixation on the healing of large allografts. J. Bone Joint Surg. 67A:657–663. 22. Capanna R., Campanacci D. A. 2001. The treatment of metastases in the appendicular skeleton. J. Bone Joint Surg. Br. 83(4):471–481. 23. Bernick S., Paule W., Ertl D., Nishimoto S. K., Nimni M. E. 1989. Cellular events associated with the induction of bone by demineralized bone. J. Orthop. Res. 7:1–11. 24. Gendler E. 1986. Perforated demineralized bone matrix: a new form of osteoinductive biomaterial. J. Biomed. Mater. Res. 20:687–697.

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25. Gendler E. 1990. Cartilage and bone induction by artificially perforated organic bone matrix. U.S. patent 4,932,973. 26. O’Donnell R. J., Deutsch T. F., Flotte T. J., Lorente C. A., Tomford W. W., Mankin H. J., Schomacker K. T. 1996. Effect of Er: YAG laser holes on osteoinduction in demineralized rat calvarial allografts. J. Orthop. Res. 14:108–113. 27. Scanlon C. E. 1991. Analysis of laser-textured, demineralized bone allografts. M.S. thesis. Northwestern University, Chicago. 28. Sires B. S. 1992. Bone allograft material and method. U.S. patent 5,112,354. 29. Nuss R. C., Fabian R. L., Sarkar R., Puliafito C. A. 1988. Infrared laser bone ablation. Lasers Surg. Med. 8:381–391. 30. Walsh Jr. J. T., Flotte T. J., Deutsch T. F. 1989. Er: YAG laser ablation of tissue: effect of pulse duration and tissue type on thermal damage. Lasers Surg. Med. 9:314–326. 31. Guo M. Z., Xia Z. S., Lin L. B. 1991. The mechanical and biological properties of demineralized cortical bone allografts in animals. J. Bone Joint Surg. Br. 73:791–794. 32. Hosny M., Sharawy M. 1985. Osteoinduction in rhesus monkeys using demineralized bone powder allografts. J. Oral Maxillofac. Surg. 43:837–844. 33. Hosny M., Sharawy M. 1985. Osteoinduction in young and old rats using demineralized bone powder allografts. J. Oral Maxillofac. Surg. 43:925–931. 34. Narang R., Wells H., Laskin D. M. 1982. Experimental osteogenesis with demineralized allogeneic bone matrix in extraskeletal sites. J. Oral Maxillofac. Surg. 40:133–141. 35. Glowacki J., Mulliken J. B. 1985. Demineralized bone implants. Clin. Plast. Surg. 12:233–241. 36. Lewandrowski K. U., Schollmeier G., Uhthoff H. K., Tomford W. W. 1998. Mechanical properties of laser-perforated and partially demineralized diaphyseal bone allografts. Clin. Orthop. 353:238–246. 37. Lewandrowski K. U., Tomford W. W., Michaud N., Flotte T. F., Schomacker K. T., Deutsch T. F. 1997. Electron microscopic studies on the process of cortical bone demineralization. Calcified Tissue Int. 61:294–297. 38. Lewandrowski K. U., Tomford W. W., Schomacker K. T., Deutsch T. F., Mankin H. J. 1997. Enhancement of incorporation of cortical bone grafts by controlled partial demineralization and laser-perforation. J. Orthop. Res. 15:748–756. 39. Lewandrowski K. U., Tomford W. W., Yeadon A., Deutsch T. F., Mankin H. J., Uhthoff H. K. 1995. Flexural rigidity in partially demineralized diaphyseal bone grafts. Clin. Orthop. 317: 254–262. 40. Lewandrowski K. U., Venugopalan V., Tomford W. W., Schomacker K. T., Mankin H. J., Deutsch T. F. 1996. Kinetics of cortical bone demineralization. A new method for modifying cortical bone allografts. J. Biomed. Mater. Res. 31:365–372. 41. Mnaymneh W, Malinin T. I., Lackman R. D., Hornicek F. J., Ghandur-Mnaymneh L. 1994. Massive distal femoral osteoarticular allografts after resection of bone tumors. Clin. Orthop. 303:103–115.

11 Material Selection for Engineering Cartilage Giuseppe M. Peretti, Jian-Wei Xu, and Mark A. Randolph Massachusetts General Hospital, Harvard Medical School, Boston, Massachusetts

I INTRODUCTION The loss or failure of an organ or tissue is one of the most critical problems in human health care, and the surgical management of structural tissue deficiency resulting from the loss, malformation, or agenesis is one of the most challenging clinical problems facing reconstructive surgeons. Advances in the medical sciences have enabled physicians to restore lost metabolic and biomechanical functions in their patients through organ transplantation, reconstructive surgery with autogenous tissue transfer, or the implantation of alloplastic materials. Although these therapies have saved or improved the lives of many patients, they remain imperfect solutions. Structural tissue deficiencies involving cartilage cause loss of function and changes in esthetic appearance, but are not generally life threatening. Unlike cases of organ failure where the life-threatening nature of the illness justifies the life-long administration of immunosuppressive therapy necessary to prevent rejection of an organ transplant, structural tissue deficiencies do not. Nonetheless, injuries to cartilage tissues adversely affect physical as well as mental health and have high economic cost [1]. Cartilage is widely distributed throughout the human body and is comprised of a combination of connective (or skeletal) tissue cells and extracellular matrix. The specific organization of the various cartilaginous tissues is related directly to the temporal and spatial functional demands on the tissue, both static and dynamic. Generally, these functional demands pertain to (1) the protection and support of related nonskeletal tissues and organs, (2) the articulations between skeletal elements, and (3) the dynamic processes related to skeletal growth [2]. Because of these differences among the cartilage tissues, the extracellular matrix, which possesses a defined biochemical composition and confers specific biomechanical properties, is represented differently among the different structures of the 195

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body. Articular cartilage is an extremely important mechanical entity in articulating joint function. For example, articular cartilage plays an important role in lubrication and wear, providing a fluid-filled, wear-resistant surface, where one diarthrodial element slides over the other in the joint [3]. Other types of cartilage tissues fulfill mechanical function as well, although different from that of the articulating surfaces of joints. Cartilage of the intervertebral disc acts as a load transmitter and shock absorber between bony vertebral bodies by transferring axial compressive forces into tangential ones, the energy of which is then absorbed and dissipated [4,5]. The functional roles of cartilage in the trachea, nose, ribs, ears, and pharynx involve maintaining form and resisting deformation, while providing some degree of flexibility [6]. There is a large population of patients who suffer from pain, stiffness, and loss of structure and function from cartilage defects resulting from burns, tumors, trauma, arthritis, or other metabolic causes. Cartilage, particularly joint cartilage, possesses limited innate ability for repair and regeneration. Consequently, injury to cartilage results in scar formation leading to permanent loss of structure and function [7,8]. In joints, for example, the articular cartilage is isolated from a vascular environment. The lack of vascularity combined with the extreme forces applied to articulating cartilage probably impede a normal wound repair process as would be observed in tissues like skin. Current surgical techniques for cartilage repair rely heavily on either autogenous composite tissue grafts or on the placement of artificial prosthetic implants. More recently, techniques have been introduced to reimplant autologous, cultured chondrocytes [9]. Each of these techniques suffers from problems which limit its clinical utility. Harvesting autologous tissue, which is limited in supply, results in donor site morbidity, is difficult to shape, and can undergo unpredictable resorption over the long term. Prosthetic metallic or plastic implants undergo migration, extrusion, and unknown long-term side effects [10]. The implantation of autologous, cultured chondrocytes into lesions in knee cartilage has demonstrated some positive clinical results, but the technique is not fully reliable in maintaining the cells in the lesion [11]. Furthermore, the repair tissue has not been fully evaluated in restoring normal biochemical composition and biomechanical properties. Finding a minimally invasive method for the in vivo generation of new cartilage to permanently repair cartilage defects in these patient populations has been the primary goal of our laboratory. One possible solution for providing quality structural tissue could be to engineer cartilage tissues to meet the requirements for the repair. As such, the material properties of synthetic or natural compounds could be manipulated to allow the delivery of an aggregate of dissociated cells into a host in a manner that will result in the formation of new functional tissue [12,13]. For example, a small sample of normal cells could be obtained from uninvolved cartilage tissues of an afflicted individual or an unrelated cadaver donor by enzymatic digestion. These free cells could then be grown in vitro to multiply to some desired number. After a sufficient quantity of cells is obtained, the cells could be combined or composited with a polymer(s) and transplanted into the defect site to restore normal function. To achieve the desired result, however, one must consider both the properties of the tissue native to the site and the properties of the polymer(s) being used to generate cartilage repair tissue. Most importantly, the engineered cartilage must integrate into the defect and heal to the surrounding cartilage at the site of implantation. Tissue engineering technology could avoid many of the potential problems associated with other forms of treatment such as reducing donor site morbidity and the potential for infection, and reducing the use of alloplastic materials. The avascular nature and relatively simple microarchitecture of cartilage makes the possibility of generating it through

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tissue engineering techniques an achievable goal. This chapter explores the biological nature of cartilage and the physical demands on cartilage from various parts of the body and discusses some polymers that have been used successfully to generate cartilage using tissue-engineering techniques. It is not the purpose of this chapter to present an exhaustive list of every polymer used for engineering cartilage or to discuss the toxicity of the various materials. Rather, the chapter presents various classes of materials and discusses their advantages and disadvantages. II CARTILAGE PROPERTIES Cartilage has several characteristics that make it particularly well suited for cell transplantation and tissue engineering. It is a relatively simple tissue in that it contains only one cell type—chondrocytes. Cartilage is a biphasic material with the solid matrix phase consisting of a dense collagen network suspended in a gel of proteoglycans (Fig. 1). The predominant collagen in articular cartilage is type II, whereas cartilage in the cranium is comprised of relatively higher amounts of type I collagen in addition to other macromolecules like elastin. Entrapped within the extracellular matrix, the chondrocytes continuously produce various macromolecules such as sulfated glycosaminoglycans to replenish the extracellular matrix. The interstitial fluid phase gives articular cartilage unique viscoelastic properties by the free flow of water and electrolytes through the proteoglycan matrix and allows for the nourishment of chondrocytes through diffusion [14,15]. The simple structure of cartilage allows for relatively pure populations of chondrocytes to be isolated from cartilage using commercially available enzymes. However, it had been consistently observed in several species that when chondrocytes are removed from their extracellular matrix they became phenotypically unstable [16–20]. That is, when chondrocytes are cultured in monolayer, the cells rapidly lose their differentiated chondrocytic phenotype by attaching to the plate, flattening, spreading, and transforming into cells characteristic of fibroblasts [16]. Several investigators have analyzed the collagen and glycosaminoglycan content in the matrix produced by these “dedifferentiated cells.” Benya et al. demonstrated that collagen type II synthesis began to decline steadily in articular chondrocytes cultured in monolayer and essentially ceased by the fifth subculture passage. In

Figure 1 Specimen of native articular cartilage from swine showing cells surrounded by cartilaginous matrix (safranin-O staining, original magnification 200).

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their studies, the final collagen product from dedifferentiated cells was 41% type I and 1% type II [19]. Others have shown that chondrocytes grown in low-density monolayer culture begin slowing the production of their characteristic large aggregating proteoglycans and begin producing predominately small nonaggregating proteoglycans [21,22]. It is clear that in monolayer culture chondrocytes assume a fibroblast-like phenotype and cease to synthesize the collagen and proteoglycan macromolecules specific to cartilage; instead the cultured cells produce macromolecules normally expressed by undifferentiated mesenchymal cells. Suspending chondrocytes in a three-dimensional matrix, similar to their natural environment, can permit the cells to retain their native phenotype and produce their extracellular components. Benya et al. reported that the culture of rabbit cartilage in agarose permitted restoration of the differentiated phenotype [22]. Others established that several culture systems including suspension culture were capable of restoring the chondrocyte phenotype [23–27]. Bonaventure et al. demonstrated that human articular cells which had dedifferentiated on monolayer culture could revert back to a chondrocyte phenotype and begin producing type II collagen and large aggregating proteoglycans in proportions characteristic of hyaline cartilage [16]. It is clear that by restoring the three-dimensional spherical shape of the chondrocyte, its characteristic phenotype can be re-expressed [28]. However, the question of whether chondrocytes taken through several cell divisions in monolayer culture can produce cartilaginous matrix of sufficient quality and quantity when used in vivo has not been determined. Chondrocytes regain their spherical shape after placing them in biodegradable hydrogel polymers allowing the cells to begin producing their characteristic matrix macromolecules [22,29,30]. Using polymers that undergo a controllable bulk erosion process in vivo, the polymer can be made to resorb at a rate proportional to the rate at which cartilaginous extracellular matrix is being deposited into the intercellular spaces. When properly orchestrated, the result is the generation of cartilage with its characteristic microarchitecture, guided by the chondrocyte’s intrinsic “programming” and facilitated by the polymer’s engineering [31]. The extracellular cartilaginous matrix plays a crucial role in the flexibility of elastic cartilage, namely, type II collagen [27] and elastin [32] fibrils. This flexibility enables elastic cartilage to withstand the tensile forces applied to the cartilage. Whereas the proteoglycans [33], which have large negative charges on them, allow cartilage to withstand compression, the high tensile stiffness and strength of collagen and elastin fibrils can sustain the surface tension put on the elastic cartilage in distraction. When the surface tension is greater than the tensile failure stress value of the collagen and elastin, a fracture can occur (Fig. 2). Recently, there has been an increased interest in cartilage fabrication for the reconstruction of craniofacial defects as a result of advances in tissue-engineering techniques [12]. Flexibility is an essential biomechanical property of cartilage in the cranium (ear and nose). Much vertebrate cartilage is covered by a specialized connective tissue called perichondrium (Fig. 3). The perichondrium forms a fibrous capsule consisting of two zones, an outer fibrous zone and inner, cellular, chondrocytic zone. The outer zone consists of dense lamellae of collagenous and elastic fibers and blends with the surrounding connective tissue. The inner zone blends smoothly with the subperichondrial cartilage. The latter is usually thicker in developing tissues, but it as absent in adult tissues [34]. Investigators have found that perichondrium has the following functions: (1) providing blood supply to cartilage [35,36], (2) regulating growth region cartilage proliferation and differentiation [37], (3) regenerating cartilage [38], and (4) cartilaginous injury repairing and intercartilaginous healing [39].

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Figure 2 Schematic diagram showing the primary components of cartilage matrix. Whereas the glycosaminoglycans are integrally involved in maintaining the cartilage under compression, when tension is applied the collagen and elastin can fracture.

Results from our initial studies on native swine cartilage demonstrated that the presence of perichondrium is also essential to the flexibility of auricular cartilage [40]. This flexibility is the result of the biomechanical properties of the cartilage itself as well as those of the adherent perichondrium and the binding of these two elements. For many years, there has been considerable interest in creating a prefabricated flexible framework for use in ear reconstruction. Current approaches for ear reconstruction [41–43], including autologous or homologous transplants and alloplastic implants, can produce wellshaped auricular frameworks, but the flexibility of the reconstructed ear is not satisfactory. Because elastic cartilage is the most important type of craniofacial cartilage, we focused on some practical methods to enhance the flexibility of tissue-engineered cartilage framework for craniofacial cartilage defect repair using tissue-engineering tech-

Figure 3 Specimen of native auricular cartilage from swine surrounded by two layers of perichondrium (hematoxylin and eosin staining, original magnification 25).

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niques. We will also discuss the possibility of providing nondegradable struts to confer favorable properties, such as flexibility and durability against fracture, to the engineered cartilage. III MATERIAL SCIENCES In the early 1970s, Greene predicted that with the advent of new synthetic biocompatible materials, transplantation of cells on such materials might result in the formation and successful engraftment of new functional tissue [20]. Although several investigators have employed natural scaffolds like collagen sponges for cartilage generation, advances in material sciences have now enabled the production of biocompatible as well as biodegradable materials that may provide favorable qualities for engineering cartilage. Langer and Vacanti postulated that if synthetic polymers could act as cell anchorage sites, they could potentially serve as three-dimensional scaffoldings for the delivery and support of transplanted cells [12]. Employing synthetic rather than naturally occurring polymers could offer many practical advantages. Chemical synthesis of polymers could allow the precise engineering of matrix configuration to permit optimal cell survival, proliferation, and subsequent tissue formation. The physical properties of synthetic matrices also can be altered in order to obtain desired characteristics of the engineered cartilage. For example, the configuration of the synthetic matrix could be manipulated to vary the surface area available for cell attachment as well as to optimize the exposure of the attached cells to nutrients. Similarly, the chemical environment surrounding a synthetic polymer might be affected in a controlled fashion as the polymer biodegrades. With a growing understanding of cell biology, it may be possible to attach ligands, such as cell-adhesion peptides, to the polymer backbone to promote or foster optimal cellular interactions. The potential also exists for binding or incorporating growth factors and hormones that could be continuously released in a controlled fashion to provide signals promoting chondrocyte differentiation and cartilage growth. The ability to add multiple side chains to the polymer structure could allow for a wide array of substances to be deployed for defined purposes. Finally, unlike the variable quality of naturally occurring polymers, chemically synthesized polymers can be consistently and reliably produced with strict quality control. A significant body of practical knowledge has emerged for manipulating synthetic polymers. Porosity has been altered to increase or decrease the intrinsic strength and elasticity of the polymer matrix as well as compressibility or creep recovery. The rate of polymer matrix degradation has been controlled by either altering the surface-to-mass ratio or by altering the surface chemistry of the polymers. Much has been learned about the chemical environment surrounding these compounds as they degrade. For instance, polyaminocarbonates cause a local basic environment and polyglycolic acid and polyanhydrides cause a local acidic environment. Polyanhydrides and polyorthoesters show surface erosion, whereas other polymers show bulk erosion. By combining polymers and altering the surface chemistries, one can favorably influence the environment into which the cells are implanted. In some instances, nondegradable materials can be utilized to enhance a particular aspect of engineered cartilage. For example, compositing or bonding engineered cartilage to nondegradable membranes to simulate the perichondrial layer could prevent fractures in the neocartilage [40]. Each of the various classes of polymers has its advantages and disadvantages with regard to cartilage tissue engineering and will be discussed in the following sections.

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IV OPEN LATTICE SCAFFOLDS A Synthetic Polymers Many early investigations for engineering cartilage from synthetic polymers focused on the use of polyesters of poly(-hydroxyesters) [12]. These polymers were chosen initially because they were already approved for human use by the Food and Drug Administration (FDA), primarily for suture material. Biodegradable polyesters such as poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), and their copolymer poly(DL-lactic-co-glycolic acid) (PLGA) possessed many desired properties to support cell transplantation (Fig. 4). These polymers can be formed into open lattice structures with high porosity that allows the free exchange of nutrients and waste products. Both PGA and PLLA have been used successfully to generate cartilage in vitro and in vivo. The lattice structure can be adjusted and provides a high surface area–to–volume ratio to allow for matrix production in the open interstices. Both PLLA and PGA degrade by hydrolysis into nontoxic metabolites and can be tailored to degrade at defined rates. Individually, PLLA, which is more hydrophobic and less crystalline than PGA, degrades slower. Polymer constructs made from PGA have been evaluated extensively by several investigators [44–47]. Constructs assessed in vitro using phase contrast microscopy demonstrate that chondrocytes adhere in multiple layers to the branching PGA polymer fibers and retain the rounded morphologic appearance typical of chondrocytes. Attachment of chondrocytes is enhanced if the polymer fibers are coated with PLLA, which is more hydrophilic. By contrast, chondrocytes grown in standard monolayer culture become flattened and lose their ability to produce differentiated cell matrix proteins. Freed and Vunjak-Novakovic have successfully grown cartilage in vitro using PGA polymer. Crucial to matrix production in vitro, however, is the placement of the constructs in a dynamic environment [48–50]. Poly(glycolic acid) polymer constructs seeded with chondrocytes have been implanted into various animal models by numerous investigators. Most studies involve implantation of cell/polymer constructs into subcutaneous pockets in nude mice. Athymic mice were selected as the predominant model because their state of compromised cell-mediated immunity allows for xenogeneic cell transplantation without immunologic rejection

Figure 4 Electron micrograph of PGA fibers seeded with sheep cartilage cells.

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of the cells. Thus, large quantitites of pure bovine calf articular chondrocytes can be easily obtained and implanted into a relatively inexpensive animal model. Histological evaluation of retrieved specimens made from bovine articular chondrocytes demonstrates tissue with the histoarchitectural appearance of normal fetal cartilage. Aldehyde fuchsin-alcian blue stains demonstrate the presence of chondroitin sulfate produced by the chondrocytes that is characteristic of cartilage matrix. Additionally, the resulting cartilage matrix demonstrates the presence of type II collagen, found almost exclusively in mammalian hyaline cartilage, and no measurable type I collagen. Kim et al. have demonstrated that implants of PGA and cells designed into specific shapes can retain these shapes during in vivo incubation in nude mice [45]. Implantation of articular chondrocytes on PGA into rabbit knee defects has demonstrated new matrix formation [44], whereas implantation of similar constructs into vascularized compartments under the skin has not resulted in cartilage [51]. These studies demonstrate that biocompatible polymer fibers seeded with chondrocytes will form new cartilage both in vitro and in vivo. Although this class of polymers is useful for generating new cartilage, these synthetic fibrous polymers have several limitations that may preclude their widespread use in cartilage repair and grafting. They are difficult to mold into desired shapes and are hydrophobic, which causes poor cellular attachment. Poor cellular attachment means there is a requirement for large numbers of cells during the cell-seeding process. This is a significant drawback since, in the clinical setting, only small biopsies of cartilage to obtain autologous cells will likely be possible, and cell numbers will have to be expanded in tissue culture. Furthermore, these polymers in the form of solid, fibrous structures require open implantation of the cell–polymer constructs since there are no methods currently available that permit these polymers to be inserted into a defect using minimally invasive means. Also, when these types of polymers have been implanted into vascular compartments (i.e., subcutaneously) in immune-competent animals, there is evidence of a strong foreign body reaction similar to that observed around polyester suture material [51]. In nude mice, a mild inflammatory response as evidenced by the presence of polymorphonuclear leukocytes and giant cells has been noted in the early time points following implantation. The resolution of the inflammatory response seems to correlate with the disappearance of the polymer fibers. B Collagen Scaffolds Collagen is the prevalent structural biomolecule in the extracellular matrix of cartilage, making it a logical choice for deriving a tissue-engineering scaffold. Collagen sponges have many desirable properties as a biological scaffold for cartilage including porosity, biodegradability, and biocompatibility. Multiple methods for preparing collagen and collagen–GAG scaffolds have been reported. Generally, collagen scaffolds are made from animal tissues such as type I collagen from bovine tendon. It is also possible to chemically modify the biomechanical and biological properties of the collagen scaffolds to enhance certain characteristics that promote cartilage formation [52]. Open lattice collagen scaffolds, some of which also include glycosaminoglycans, have been synthesized and used for generating new cartilage matrix. Scaffolds made from a single collagen type or composites of two or more types have been employed. All seem to show favorable results with regard to chondrocyte adherence and their ability to maintain a differentiated chondrocyte phenotype [53,54]. In comparison to other open lattice

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synthetic scaffolds, it appears that collagen sponges promote collagen production, whereas synthetics such as PGA promote proteoglycan synthesis [55]. Additionally, the porosity of the sponge and the surface characteristics of the meshwork are critical elements to consider for promoting cell growth and cartilage formation. Some characteristics, such as pore size, can be changed by varying the freeze-drying process [54]. Collagen sponges can be also be modified using growth factors or other manipulations to promote chondrocyte growth and cartilage matrix formation. For example, sponges can be impregnated with exogenous growth factors that promote matrix production such as bFGF [56]. Studies from Mizuno and Glowacki have shown that collagen sponges can also incorporate other components that induce differentiation of nonchondrocyte cells to differentiate and form cartilage. For instance, those studies demonstrated that the addition of demineralized bone powder to the collagen sponges would stimulate human dermal fibroblasts to synthesize an extracellular matrix resembling that of the cartilaginous tissue [57]. Despite their appeal as a biological material for making a tissue-engineering scaffold, collagen sponges have some disadvantages. These scaffolds can cause a foreign body reaction with a thin fibrous capsular tissue surrounding collagen sponge implants [56]. Such reactions may interfere with the integration of new cartilage formed in the scaffold with the surrounding native cartilage in the recipient site. C Hyaluronan-Based Scaffolds Hyaluronan is one of the major constituents of undifferentiated mesenchyme in the developing embryo as well as in the cartilage extracellular matrix. Hyaluronan has been shown to support proliferation of mesenchymal progenitor cells and differentiation into chondrocytes. Additionally, in cartilage it is believed to play a significant role in physical microenvironments affecting chondrocyte function [58–60]. It is also possible that hyaluronan creates an environment during cartilage repair that recapitulates processes during embryonic development. Hyaluronan can be formulated into many different chemical and physical entities that provide a favorable environment for cartilage generation, allowing both synthesis of matrix components and differentiation of the progenitor cells [61,62]. HYAFF-11® is made of a linear derivative of hyaluronan modified by complete esterification of the carboxylic function of the glucuronic acid moiety with benzyl groups [63]. HYAFF-7 ® is similar except for being an ethyl ester instead of a benzyl ester. The fiber thickness of both types is in the range of 20 m and processed into nonwoven meshes. HYAFF-7 degrades approximately 30 days in vitro and 60 days in vivo, whereas HYAFF-11 degrades somewhat slower, 60 days in vitro and approximately 110 days in vivo [61]. Another crosslinked hyaluronan formulation generated by a condensation procedure has been reported by Solchagá et al. In this formulation, the polymer was stabilized by directly esterifying some of the carboxylic groups of glucuronic acid along the chain with hydroxyl groups of the same or different hyaluronan molecules [63]. The resulting polymer has a porosity of 85% with pores of 10–300 m. The recent study from Solchogá et al. showed encouraging results in vivo in the treatment of an osteochondral defect in a rabbit knee model using hyaluronan-based polymer. The main hypothesis of their study was that hyaluronan fragments could encourage migration of the mesenchymal stem cells from the bone marrow to a chondrogenic differentiation and repair the osteochondral defect in the rabbit condyle. Results showed a higher

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reparative rate in the lesion repaired by the polymer sponges than those untreated or those treated with HYAFF-11 [63]. D Cartilaginous Scaffold The use of other biological material as scaffold for delivering chondrocytes has been investigated in our laboratory. Peretti et al. [64] reported on the capacity of devitalized cartilage matrix to serve as a scaffold. They reported on the capacity of isolated chondrocytes seeded onto devitalized cartilage matrix to bond the pieces of the matrix together in vivo. Devitalized lamb articular cartilage was either co-cultured with viable allogeneic chondrocytes (experimental) or cultured without cells (control). The co-culturing phase allowed the isolated chondrocytes to adhere to the surfaces of the devitalized matrix. Composites of three such slices were constructed with fibrin glue and implanted into nude mice up to 42 days. Results demonstrated that bonding of the experimental matrices with viable chondrocytes was achieved at 28 and 42 days (Fig. 5), whereas no bonding occurred in control composites without viable chondrocytes. We concluded that devitalized cartilage matrix could serve as a scaffold to which isolated chondrocytes could attach and repopulate, and confirmed that the chondrocytes could produce matrix that bonded or healed the cartilage slices together. In subsequent studies [65] this integration between the newly formed cartilaginous tissue and the devitalized cartilage matrix scaffold was better characterized and quantified by biomechanical analysis of the neocartilage generated between devitalized sections. Devitalized cartilage discs were incubated in the presence (experimental) or absence (control) of isolated chondrocytes in suspension culture in order to allow chondrocytes to adhere to the dead matrix. Cartilage discs were then held in apposition and implanted subcutaneously in nude mice up to 6 weeks. Mechanical testing was performed on the samples by tensile testing on a Dynastat mechanical spectrometer. The results showed that tensile strength, fracture strain, fracture energy, and tensile modulus increased with time for experimental group while a control group showed no increase (Fig. 6). These studies have led to the development of a mouse model for repairing the avascular portion of knee meniscus tissue [66]. Devitalized meniscus specimens measuring

Figure 5 Neocartilage (NEO) synthesized between two devitalized cartilage specimens (DC) by articular chondrocytes, which were seeded onto the dead matrix surfaces (safranin-O staining, original magnification 200). (From Ref. 64.)

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Figure 6 Time course of changes in (A) tensile strength, (B) fracture strain, (C) fracture energy, and (D) tensile modulus over 6 weeks in vivo for constructs seeded with chondrocytes (and unseeded controls) and implanted in pairs in nude mice. All data are shown as mean SD and represent the biomechanical properties of the bonding tissue formed between two devitalized cartilage discs. (From Ref. 65.)

4  2  0.5 mm were derived from lamb knee menisci and were co-cultured with viable allogeneic lamb articular chondrocytes. A 4-mm bucket-handle lesion was made in the inner third of other devitalized lamb menisci, and the tissue-engineered composite was sutured inside the lesion. Samples were implanted in subcutaneous pouches of nude mice for 14 weeks. Similar control composites were made from menisci treated with nonviable cartilage chips that had not been co-cultured with viable chondrocytes; other control groups consisted of torn menisci treated only with suture or with the lesion left untreated. Results showed that seven out of eight experimental samples having cells delivered on the scaffold achieved repair of the bucket-handle lesion, whereas none of the control groups demonstrated repair. Histological analysis confirmed notable bonding between newly formed tissue from transplanted cells and the edges of the meniscus fracture. These results show that biological materials can be used as scaffold to deliver cells and effect repair of cartilage lesions. V HYDROGEL SCAFFOLDS The inability to deliver chondrocytes through minimally invasive techniques when using fibrous or open lattice type polymers stimulated investigations into other types of polymer carriers such as hydrogels. Hydrogels are gelatinous colloids that when maintained under controlled conditions exhibit three-dimensional stability. Hydrogels are produced by mixing a soluble polymer in water and adding a crosslinking agent to gel the mixture. As the polymer gels, there is usually sufficient opportunity to mold and shape the final three-dimensional configuration of the hydrogel. Additionally, by existing in a liquid phase, these

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polymers have the potential for injectable delivery. The combined high water content and elasticity of polymer hydrogels lead to many tissue-like properties of these materials, making them ideal candidates for tissue-engineering matrices. Hydrogels have proven to be extremely effective in providing a hospitable, three-dimensional support matrix for the immobilization of cells. By suspending chondrocytes in a highly porous aqueous matrix, they can maintain their differentiated function and are capable of producing large quantities of extracellular matrix macromolecules [30,67]. It has been hypothesized that chondrocytes will provoke an immune response when their cell membranes are exposed to the immune system after being removed from their protective matrix [68]. If the chondrocytes of higher mammals do elicit an autoimmune reaction when exposed, polymers such as polyglycolic acid with an open latticework of fibers may be less desirable as a carrier since immune cells would have unencumbered access to the transplanted chondrocytes. If there is, indeed, an in vivo immune response to the chondrocytes of higher mammals, biocompatible hydrogel polymers may attenuate this immune response until new matrix is formed. As the field of tissue engineered cartilage moves forward into preclinical testing in immune-competent animals and closer to clinical application, the role that the carrier polymer matrix plays must be elucidated for any polymers under consideration as well. Additionally, the cells may be shielded from the immune system long enough for them to produce their own protective matrix. Examples of hydrogels used to encapsulate cells include ionically crosslinked alginates [30,69–71] and chitosan [72,73], hydrogen-bonded block copolymers such as Pluronics [74], and covalently crosslinked fibrin glue. A Natural Hydrogels 1 Alginate This hydrogel has become the most widely used hydrogel polymer for encapsulating living cells. Alginate, extracted from brown seaweed algae, is a polysaccharide consisting of a heterogeneous group of linear binary block copolymers of D-mannuronic acid and L-guluronic acid residues. Each type of algae produces alginate with different gel-forming capacity, porosity, and strength. Gel formation occurs when two consecutive L-guluronic acid residues form binding sites for calcium ions and long sequences of such sites produce crosslinks with similar arrangements on other alginate molecules. The result is an open lattice of polymerized polysaccharide molecules. Consequently, the gel-forming capacity of a given batch of alginate is dependent upon the content of L-guluronic acid and the presence of free calcium ions. Similarly, the porosity and the diffusion characteristics of a given alginate gel are also dependent upon the content of L-guluronic acid. Using electron microscopy, the pore size of 2% calcium alginate gel has been found to range from 5 to 200 nm. Thus, the diffusion of small molecules such as glucose is little affected by L-guluronic acid content, while the diffusion of larger molecules such as proteins are significantly increased when the L-guluronic acid content is higher. With high L-guluronic acid content, even large proteins such as fibrinogen with molecular weights greater than 68 kDa are capable of diffusion. High L-guluronic acid content also affects mechanical strength of the gel. For example, gels with greater than 70% L-guluronic acid produce the strongest polymers. Higher molecular weights related to longer polymer chain length also significantly improve strength [75]. The immobilization of chondrocytes in alginate was first described by Gou et al. in 1989 [76]. They demonstrated that entrapped chondrocytes maintained their differentiated

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phenotype and were capable of proliferating at rates approximately 30% of that achieved in monolayer culture. Later, Ramdi et al. and Hausselman et al. independently showed that chondrocytes maintained production of sulfated glycosaminoglycans and type II collagen even when encapsulated for up to 8 months in calcium alginate gel [67,77]. Unlike in monolayer culture where newly synthesized proteoglycans and collagens diffuse into the surrounding media, alginate was found to contain these macromolecules in the form of an intact interterritorial matrix. Using a different hydrogel, Bushmann was able to demonstrate that entrapped chondrocytes, with their newly synthesized macromolecules, formed a mechanically functional matrix with increasing stiffness and compression strength during long-term culture [21]. Paige et al. embarked on a series of studies examining the feasibility of using calcium alginate as a vehicle for chondrocyte transplantation in the form of injectable cartilage [70]. They postulated that since alginate is degraded by enzymatic pathways into its two monomeric units and excreted in the urine [69], is nontoxic, and invokes a minimal immune response [69,78], it could potentially serve as a temporary three-dimensional scaffolding to support the chondrocytes until a new matrix is formed. In their initial studies, they demonstrated the production of cartilage matrix in vivo by entrapping bovine articular chondrocytes in alginate gel and implanting the polymerized constructs into athymic nude mice [71]. Freshly isolated bovine articular chondrocytes were mixed with a 1.5% sodium alginate solution in cell densities of 0, 1, 5, and 10 million cells/mL. Discs of 125 L of cell–alginate solution were molded in vitro by contact with a polymerizing CaCl solution. The discs were then implanted subcutaneously on the dorsum of athymic nude mice and harvested after 8 and 12 weeks. At harvest, the implants were found to be surrounded by a thin capsule invested with a thin capillary network and only a minimal inflammatory reaction. The gross appearance of specimens resembled that of cartilage in constructs with original cellular densities of 5  106 cells/mL and higher, but not in those at lower cell density. This same pattern was seen on histological analysis where constructs having higher original encapsulation cell densities revealed a histoarchitecture resembling hyaline cartilage with isolated cells aggregated in nodules of basophilic ground substance. Highly sulfated proteoglycans and collagen were demonstrated in the pericellular matrix with special stains. Biomechanical analysis demonstrated compression strengths 11 times stronger in constructs starting with 10 million cells/mL over acellular controls of alginate alone. The mean compression strength of specimens as well as their mean weight were found to generally increase with increasing original cellular densities and periods of in vivo implantation. This study demonstrated that calcium alginate could be used as a temporary three-dimensional support matrix for the de novo generation of cartilage. However, since alginate polymerizes almost immediately on contact with free calcium ions, the chondrocyte/alginate constructs had to be polymerized in vitro before being implanted in an open procedure. Seeking to create an injectable construct, Paige and colleagues investigated the use of CaSO4 as the polymerizing agent [70]. Unlike CaCl, which easily dissociates liberating free calcium ions, CaSO4 dissolves poorly as SO4 has a higher affinity for calcium then does Cl, thus liberating free calcium more slowly. Employing this property of CaSO4 to slow down the polymerization of the cell/alginate constructs allows time for injection and manipulation prior to gelation. Using these methods they were able to create an injectable cell/polymer construct for the creation of new cartilage in vivo. Routine histological analysis was augmented with immunohistochemistry to reveal the presence of type II collagen both pericellularly and intracellularly.

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Although alginate has been used successfully to generate cartilage in vivo, the biological degradation of alginate in vivo is not known. The cartilage matrix produced in these studies was nonuniform and demonstrated only islets of cartilage rather than contiguous cartilage matrix throughout. Thus, the quality of the cartilage is indeterminate, and it is not known if the alginate will fully resorb or break down over time in vivo. 2 Fibrin Glue Polymer Fibrin glue polymer has been studied extensively in our laboratory and by others as a scaffold for engineering cartilage from isolated chondrocytes. Clinically, fibrin glue can be made from autologous blood products. Sims et al. demonstrated that articular chondrocytes encapsulated in fibrin polymer made from human blood cryoprecipitate permitted neocartilage formation when cell/polymer constructs were implanted into nude mice [79]. Silverman et al. investigated the feasibility of using a fibrin glue polymer to produce injectable tissue-engineered cartilage. This series of experiments focused on determining the optimal fibrinogen and chondrocyte concentrations required to produce homogeneous cartilage matrix in nude mice [80]. The study demonstrated that the most favorable concentration of fibrinogen was 80 mg/cc by measuring the rate of degradation of fibrin glue using varying concentrations of purified porcine fibrinogen. The fibrinogen was mixed with thrombin (50 units/cc in 40 mM CaCl) to produce fibrin glue. Swine chondrocytes were suspended in the fibrinogen prior to the addition of thrombin. The chondrocyte/polymer constructs were then injected into the subcutaneous tissue of nude mice using chondrocyte densities of 10, 25, and 40 million chondrocytes/cc of polymer (0.4-cc injections). At 6 and 12 weeks the neocartilage was harvested and analyzed by histology, mass, glycosaminoglycan (GAG) content, DNA content, and collagen type II content. These studies demonstrated that fibrin glue is a suitable polymer for the formation of injectable tissue engineered cartilage in the nude mouse model and that a concentration of 40 million chondrocytes/cc yielded the best quality cartilage at 6 and 12 weeks (Fig. 7). Fibrin glue gel has also been useful for encapsulating cells for in vitro studies [81]. Utilizing a similar in vitro model, Fortier et al. investigated the behavior of equine chondrocytes seeded into fibrin glue discs and cultured in the presence of exogenous insulinlike growth factor-1 (IGF-1). Chondrocyte/fibrin constructs were cultured for 14 days, with IGF-1 added to the culture medium at varying concentrations. This study demonstrated the

Figure 7 Histological analysis of tissue engineered cartilage and native cartilage. (A) Native articular cartilage. (B) Tissue engineered cartilage made of fibrin glue and articular chondrocytes after 6 weeks and (C) 12 weeks in vivo in nude mice (safranin-O staining, original magnification 100).

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Figure 8 Neocartilage (NEO) sandwiched between darker stained native cartilage (NC) (toluidine blue staining, original magnification 100). (From Ref. 83.) ability of articular chondrocytes in maintaining their phenotype during the three-dimensional culture in fibrin glue. Moreover, the data showed that the supplementation with 50 and 100 ng of growth factor per milliliter enhanced the synthesis of matrix products by the isolated chondrocytes [82]. To study the healing capability of cartilage engineered with fibrin polymer, Silverman et al. performed experiments analyzing the interaction between the engineered cartilage and native cartilage in vivo [83]. Tissue-engineered cartilage was produced using fresh swine articular chondrocytes (40  106 cells/cc) suspended in a fibrin glue polymer. The polymer/chondrocyte mixture was sandwiched between two discs, 6 mm in diameter, of fresh swine articular cartilage. These constructs were implanted into a subcutaneous pocket on the backs of nude mice. The constructs were harvested 6 weeks later and assessed histologically, biomechanically, and by electron microscopy. Control samples consisted of cartilage discs held together by fibrin glue alone (no chondrocytes). Histological and electron microscopic evaluation of the experimental constructs revealed a layer of neocartilage between the two cartilage discs (Fig. 8). The neo-cartilage appeared to fill all irregularities along the cartilage disc surface without any gaps. Safranin-O and toluidine blue staining indicated the presence glycosaminoglycans and collagen, respectively. Control samples without cells showed no evidence of neocartilage formation despite the close apposition of the viable cartilage discs. The mechanical properties of the bonded experimental constructs, as calculated from stress–strain curves, differed significantly from that of the control samples (Table 1). Failure was observed to occur in all cases at the interface between the neocartilage and the native cartilage. This study demonstrated that tissue-engineered Table 1 Biomechanical Properties of the Bonded Experimental Constructs, as Calculated from Stress–Strain Curves

Modulus, MPa Tensile strength, MPa Failure strain, % Fracture energy, J Source: Ref. 83.

Constructs (mean)

Controls (mean)

P value

0.74 0.065 0.17 0.00049

0.21 0.0098 0.040 3.1E-05

0.000108 0.000185 0.000178 0.003584

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cartilage produced using a fibrin-based polymer permitted adherence to adjacent cartilage and could withstand forces significantly greater than those cartilage samples adhered only by fibrin glue. The results of all these studies demonstrate that fibrin glue polymer has many advantages as a scaffold for engineering cartilage tissue. Fibrin glue allows cell distribution in a three-dimensional orientation permitting the chondrocytes to maintain their characteristic phenotype and to produce cartilaginous extracellular matrix, critical elements for the biomechanical characteristics of this tissue. Moreover, it is biocompatible, and the biodegradation can be controlled somewhat using agents like aprotinin or aminocaproic acid that prevent degradation of the fibrinogen [80,84]. However, Silverman et al. noted a significant reduction in volume after implantation of swine articular chondrocytes into nude mice [80]. The final volume of the constructs at the time of explantation was only about 30% of the original implant volume. To avoid this volume reduction of the scaffold, probably intrinsic to the use of fibrin glue, Peretti et al. performed studies on adding devitalized cartilage chips to the fibrin glue/chondrocyte mixture. They postulated that addition of these matrix chips could enhance the mechanical properties of the reparative tissue as well as prevent volume reduction [85]. Lyophilized articular cartilage chips measuring between 500 to 1000 m were mixed with a cell/fibrinogen solution and thrombin to obtain constructs made of fibrin glue, chondrocytes, and devitalized cartilage chips. The data from this study demonstrated that adding devitalized cartilage matrix to the isolated chondrocytes in fibrin glue permitted new cartilage matrix formation without a loss of volume (Fig. 9). Subsequent biomechanical analyses of similar constructs showed higher modulus and lower hydraulic permeability values in these experimental samples with respect to the other groups [86,87]. The results from that study demonstrated that autologous chondrocytes transplanted in a composite with fibrin glue and devitalized cartilage chips produce a tissue with cartilaginous appearance and biomechanical integrity. Importantly, the study confirmed that these composites have the capacity of maintain their original mass. To further study the relationship between the volume reduction of fibrin glue polymer constructs using articular chondrocytes, Xu et al. performed studies using auricular and

Figure 9 Macroscopic view of the composites at 9 weeks following implantation in nude mice. The picture shows the differences in size of the specimens belonging to the different study groups: Group A: chondrocytes  fibrin glue  cartilage chips; Group B: fibrin glue  cartilage chips; Group C: chondrocytes  fibrin glue; Group D: fibrin glue alone. (From Ref. 85.)

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costal chondrocytes in the gel scaffold. At the same cell density (40  106 cells/mL) that Silverman used [80], auricular and costal chondrocytes were suspended in fibrin glue and the constructs were implanted into subcutaneous pouches in nude mice for 4, 8, and 12 weeks. Ratios between final and initial weight and dimensions were calculated. The ratios of constructs made from auricular chondrocytes increased by 20–30% by the latest time point. The ratios of the costal chondrocyte group decreased in the early stages after implantation. However, the final ratios at 12 weeks were restored to close to the initial values. These studies demonstrated that the balance between the absorption of polymer scaffold and production of cartilaginous matrix is another key point to control the volume of the chondrocyte–fibrin glue constructs. This suggests that auricular and costal chondrocytes are preferable to articular chondrocytes in their ability to synthesize a matrix that maintains the implant volume in the subcutaneous environment [88]. B Synthetic Hydrogel Polymers 1 Poly(Ethylene Oxide) Sims et al. and Elisseeff et al. have explored the use of poly(ethylene oxide) (PEO) in an effort to overcome some of the limitations of biological hydrogels [89,90]. Poly(ethylene oxide) is a linear polyether with repeating molecular units of (CH2CH2O)n. Unlike calcium alginate, which is polymerized by the addition of a crosslinking agent, poly(ethylene oxide)’s viscosity is increased by increasing its molecular weight and concentration. These variables can be adjusted to provide the viscosity of choice, which remains stable throughout the preparation and injection steps. This allows greater control in delivering and shaping the implant both during and after injection. When placed in vivo, PEO is resorbed by 6 weeks and excreted through the kidneys [91,92]. Poly(ethylene oxide) possesses certain other characteristics that make it a particularly interesting hydrogel for use in cellular transplantation. It is both hydrophilic and hydrophobic and, therefore, is soluble in water as well as organic substances. While it can interact with cell membranes and some polyanions, it has a tendency to exclude other substances, and, consequently, it is virtually noninteractive with biological molecules. As a result, it has been used extensively in medical devices to decrease protein and platelet deposition. Interestingly, poly(ethylene oxide) can be modified to interact favorably with cells. Ligands such as cell-adhesion peptides can be linked to the free terminal hydroxyl group to foster optimal cellular interactions. Hubbel [93] has shown that endothelial cells can be made to preferentially adhere to vascular grafts lined with poly(ethylene oxide) bound to specific ligands. The potential also exists for binding growth factors to promote growth or even guide differentiation. To test the ability of poly(ethylene oxide) to serve as a temporary support vehicle for the transplantation of chondrocytes, Sims mixed freshly harvested bovine articular chondrocyte in a solution of 20% PEO (MW 100,000) for a final cellular concentration of 10 million cells/mL [89]. The cell/polymer mixture was injected subcutaneously in athymic nude mice and allowed to incubate for 6 and 12 weeks. Within the first week of transplantation, the implants had the consistency of soft moldable putty, which progressed to firm nodules by 3 weeks. At harvest, the implants demonstrated cartilage formation both histologically and biochemically. Although these studies demonstrate that PEO is suitable as a support matrix for the generation of new cartilage, it is not an ideal vehicle for injection and subsequent molding. Since PEO is not polymerized into discrete shapes after injection, but rather is injected at its maximum viscosity, it remains in this state until new cartilaginous

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matrix begins to replace it and add stiffness. Consequently, it does not maintain a given configuration once molded and, therefore, remains a suboptimal support agent for injectable cartilage. Poly(ethylene oxide) molecules can be crosslinked by chemical reaction or radiation to produce hydrogels that retain shape. Hill-West et al. has developed a method by which a photosensitive initiator added to the end groups of PEO can form crosslinks between molecules by activating the initiator groups with UV light [94,95]. This process has been used to polymerize PEO in situ. Elisseeff demonstrated that chondrocyte/PEO constructs can be injected subcutaneously, molded to the desired shape, and then polymerized transdermally with ultraviolet light [90,96]. Poly(vinyl alcohol) (PVA), another photocrosslinkable hydrogel, and poly(ethylene oxide) were chosen for the macromer backbone because of their long history in medical applications and desirable chemistry allowing easy modification. These polymers and gelation process are designed to provide easy placement (through photocrosslinking chemistry); to provide mechanical and structural stability with desirable transport properties during the regeneration process (through chemical modifications and photografting); and to allow the formation of complex shapes with suitable adhesion to treat craniofacial and joint defects. Rational design of these types of materials could optimize the immunoprotective capacity of the photopolymerizeable gels through modifications of the network structure and chemistry. 2 Pluronic Pluronic is a block copolymer made from combinations of poly(ethylene oxide) and poly(propylene oxide) (PPO). Ethylene oxide provides the hydrophilic component, while propylene oxide contributes the hydrophobic component. As such, the PPO component can interact with lipids in the cell membranes, while the PEO component interacts with the aqueous extracellular milieu. The percentage of either component, as well as the total number of each, can be varied to provide a final polymer with the desired properties. Like pure PEO, the pluronic molecules can be conjugated to proteins to foster cellular interactions. Pluronic was introduced in the 1950s and has been used extensively in the chemical, agricultural, and pharmaceutical industry [97]. It has been used to coat the surfaces of medical devices to reduce the degree of protein absorption and platelet adhesion and has been found to be nontoxic. Of particular interest to cartilage tissue engineering is the gel-forming capacity pluronic. It is thought to form gels by the hydrogen bonding that occurs in aqueous phase between the ether oxygen groups of pluronics and the protons in water. Thus, its gelling efficiency can be improved by increasing the level of oxyethylation and the molecular weight. Additionally, gel strength can be supplemented by adding agents such as sucrose (5%w/w) whose hydroxyl groups enhance hydrogen bonding. It is a liquid at low and high temperatures but forms gels at room or body temperatures. This property allows cells to be mixed at 4°C with liquid pluronics and forms a gel when placed in vivo at body temperature. This allows some degree of control over shape when injecting chondrocyte/polymer constructs. Studies in our laboratory and others have demonstrated the use of pluronics for generating cartilage. Ashiku et al. investigated the use of pluronics F-68 to encapsulate rabbit auricular and articular chondrocytes for injection as cell/polymer constructs into nude mice [98]. Using this method, the specimens produced had the pearly opalescence and consistency of cartilage, and on standard hematoxylin and eosin stains demonstrated a histoar-

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chitecture consistent with cartilage. Safranin-O stains of specimens confirmed the presence of proteoglycans. Cao et al. and Arevalo-Silva et al. have used pluronic F-127 to generate cartilage in an autologous swine model with favorable results [51,99]. Although there have been several reports using pluronics for generating cartilage, little is published on the resorption rate of the polymer in vivo when generating cartilage in situ. VI INERT NONRESORBABLE MATERIALS Investigators have demonstrated numerous techniques for improving the biological and biomechanical properties of tissue-engineered cartilage. Some key strategies for improving these properties include techniques (1) to improve the bioproperties of extracellular cartilaginous matrix, (2) to provide internal support to tissue-engineered cartilage, and (3) to add external (pseudoperichondrium) support to tissue-engineered cartilage. One objective is to improve the flexibility of the tissue-engineered cartilage framework, particularly in cartilage tissues other than articulating joint cartilage. This section discusses some possibilities to enhance the neocartilage matrix properties by incorporating nonresorbable materials to meet the needs for reconstruction. Cao et al. engineered cartilage in the shape of human auricles in nude mice using articular chondrocytes and a biodegradable internal PGA/PLLA scaffold to attain the desired shape of an ear [47]. Although their work demonstrated proof of principle for the technique, using auricular cells instead of articular cells could have improved their results. Nonetheless, the investigators demonstrated that tissue-engineered cartilage could be generated with a resorbable endoskeletal scaffold. There have been no reports of successfully generating complex three-dimensional cartilage structures using PGA in immune-competent animals, however. This may be due to limitations of these types of polymers, which are subjected to inflammatory responses. To avoid the inflammatory response toward PGA type scaffolds, Arevalo-Silva and colleagues [99] investigated the use of nonbiodegradable endoskeletal scaffolds made from the following materials: (1) high-density polyethylene, (2) soft acrylic, (3) polymethylmethacrylate, (4) extrapurified silastic, and (5) conventional silastic. They concluded that using a permanent biocompatible endoskeleton demonstrated success in limiting the inflammatory response to the scaffold, especially the high-density polyethylene, acrylic, and extrapurified silastic. Despite the success of using these materials, the investigators did not comment on the flexibility of the constructs and other biomechanical properties of neocartilage. Studies in our laboratory examined the mechanical function of perichondrium in applying external support to ear cartilage. We found that intact perichondrium prevented fracture of ear cartilage tested. From these studies we concluded that providing a perichondrial layer was important to confer flexibility to engineered cartilage tissue intended for craniofacial reconstruction. To simulate a perichondrial layer, we investigated expanded polytetrafluoroethylene and lyophilized perichondrium as structural components for supporting the engineered cartilage. A Expanded Polytetrafluoroethylene Expanded polytetrafluoroethylene (ePTFE) is a biocompatible material that has been used successfully in a multitude of biomedical and clinical applications [100–106]. One advantage of this material is its microporous structure, which allows biointegration for soft tissue fixation as well as overall mechanical integrity [107]. The results from the gross mechani-

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cal testing (Fig. 10) on engineered cartilage ePTFE composites demonstrated that ePTFE membrane matched the needs for creating a pseudoperichondrium for engineered cartilage. The ePTFE membrane is firm enough to sustain the tension placed on the surface of the engineered cartilage ePTFE composite. The adhesive character of fibrin polymer allowed the fibrin glue chondrocyte composite to combine with ePTFE membrane compactly at the outset of the experiment. Subsequently, the chondrocytes permeated into the micropores of ePTFE membrane and produced neocartilaginous matrix forming a tight bond between ePTFE membrane and engineered cartilage (Fig. 11). This integration of cartilage and ePTFE membrane formed a flexible cartilage framework with a pseudoperichondrium. Gross mechanical testing of two experimentally engineered cartilage groups demonstrated that the ePTFE membrane could not provide enough mechanical support to engi-

Figure 10 (A) Gross mechanical testing on the group in which ePTFE was put in the center of the constructs. The specimens fractured when they were bent in one (a) or in opposite direction (b). The composites were destroyed by torsion testing (c). (B) When the ePTFE was placed on both surfaces of composites, they did not fracture by bending in either direction (a). The specimens could recover after bending test (b). The specimens could not be destroyed by torsion test (c), and the samples recovered from this testing as well (d).

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Figure 11 Bonding between ePTFE and cartilaginous matrix: The transplanted chondrocytes penetrated into the microporous structure of the ePTFE (→ ←) and manufactured neocartilaginous matrix. The newly formed cartilaginous tissue formed a tight bond to the ePTFE membrane (hematoxylin and eosin staining, original magnification 50). (From Ref. 108.)

neered cartilage when it was placed in the middle of engineered cartilage ePTFE composite. Only when the polymer was placed on both the surfaces of the composite did the ePTFE membrane maintain flexibility of tissue-engineered cartilage. Thus, recreating a pseudoperichondrial layer similar in structure and position to that of native perichondrium could provide the necessary flexibility for making suitable tissue-engineered cartilage for craniofacial repair [108]. B Lyophilized Perichondrium Although ePTFE adds flexibility, being an alloplastic material might cause a foreign body reaction and limit its clinical application [41]. Autologous perichondrium would be preferable for making new perichondrium for engineered cartilage. However, an insufficient supply of autologous perichondrium limits clinical applicability. We propose using a xenogeneically derived material, lyophilized swine auricular perichondrium, to serve as a pseudoperichodrium for generating a flexible tissue-engineered cartilage framework. Pilot studies demonstrated that lyophilized perichondrium could bond tightly with chondrocytes–fibrin polymer constructs in nude mice. The results from gross mechanical testing revealed that lyophilized perichondrium can confer flexibility to tissue-engineered cartilage in a similar fashion to ePTFE membranes. Other natural and synthetic materials could serve to enhance the mechanical function of engineered cartilage. VII OTHER CONSIDERATIONS FOR ENGINEERING CARTILAGE Engineering cartilage relies heavily on the cells to produce extracellular matrix products and the scaffold material selected. There are other considerations for generating cartilage as well. If cartilage tissue is to be generated entirely in vitro, the culture conditions play a significant role in seeding the polymers and the composition of the final cartilage matrix. Additionally, growth factors have been shown to have a significant effect on the proliferation of cells and chondrogenic differentiation. These growth factors can be added to cell cultures or incorporated into the cell polymer scaffolds.

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A Hydrodynamic Cultivation Conditions It was demonstrated that hydrodynamic conditions in tissue culture bioreactors can modulate the composition, morphology, mechanical properties, and electromechanical function of tissue-engineered cartilage. The relationships between the composition and mechanical properties of engineered cartilage constructs have been studied by culturing bovine calf articular chondrocytes on fibrous PGA scaffolds in three different environments: (1) static flasks, (2) mixing flasks, and (3) rotating vessels. After 6 weeks of cultivation, the composition, morphology, and mechanical function of the constructs in radially confined static and dynamic compression all depended on the conditions of in vitro cultivation. Static culture yielded small and fragile constructs, while turbulent flow in mixed flasks yielded constructs with fibrous outer capsules. Both environments resulted in constructs with poor mechanical properties. The constructs that were cultured freely suspended in a dynamic laminar flow field in rotating vessels were the largest, contained continuous cartilage-like extracellular matrices with highest fractions of glycosaminoglycan and collagen, and were not significantly different from natural cartilage explants [50]. The effects of mixing on the composition of engineered cartilage can be compared with other reported effects of hydrodynamic forces on cultured chondrocytes and cartilage explants. Steady hydrodynamic shear enhanced glycosaminoglycan synthesis in chondrocyte monolayers [109], whereas intermittent motion of medium in roller bottles stimulated chondrocytes to form cartilaginous nodules [110]. Mixing during cell seeding and tissue cultivation resulted in increased fractions of tissue components in three-dimensional cartilaginous constructs [49]. In particular, the constructs grown in rotating vessels had a continuous cartilaginous matrix with glycosaminoglycan fractions ranging from approximately 5% wet weight after 6 weeks of cultivation [48] to approximately 8.8% wet weight after 7 months of cultivation [111]. Constructs cultured in rotating vessels were uniformly cartilaginous throughout their entire cross-sections and had a very high, stable fraction of type II collagen (92–99% of the total collagen of the construct [48]. It is possible that dynamic changes in hydrodynamic forces acting on freely settling constructs in rotating flow resemble some aspects of the dynamic mechanical loading, which enhanced glycosaminoglycan synthesis in cartilage explants [112], in contrast to static loading or the absence of loading, which caused cartilage matrix degradation [113]. B Exogenous Growth Factors Investigators have demonstrated that growth factors can not only stimulate chondrocyte expansion in a short period of time, but also enhance chondrocytes to produce useful cartilaginous extracellular matrix. Basic fibroblast growth factor demonstrated a positive influence on the in vitro and in vivo growth of engineered human auricular cartilage [114]. Chondrocyte synthesis of type II collagen is most strongly stimulated by insulinlike growth factor-1 [115]. Transforming growth factor- (TGF-) and IGF-1 stimulate glycosaminoglycan synthesis [116], whereas exogenous growth factors can improve the quality of engineering cartilaginous tissue in vitro. Growth factors have a potential role in clinical application where the goal will be to generate a large volume of high-quality tissueengineered cartilage from a small donor specimen in a short period of time. C Endogenous Growth Factors Gene transfection techniques can be used for engineering cartilage by transfecting cells with growth factor genes to overstimulate synthesis of cartilaginous extracellular matrix.

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Investigators have demonstrated that some growth factors could stimulate chondrocyte synthesis of type II collagen and proteoglycans, which are the crucial elements for maintaining the bioproperties of tissue-engineered cartilage. Transfer of growth factor genes, such as IGF-1 gene, TGF- gene, and bone morphogenetic protein-2 (BMP-2) gene to chondrocytes could greatly increase matrix synthesis in vitro [115]. Moreover, cells from infected cultures maintain a chondrocytic phenotype and continue to express elevated growth factor [117]. Mi et al. has shown that it is possible to transfer human IGF-1 gene into rabbit knee joints by a first-generation adenoviral vector to promote proteoglycan synthesis without significantly affecting inflammation or cartilage breakdown [118]. Although their data showed increased extracellular matrix production, the investigators did not show any biomechanical data on the newly formed cartilage compared with native cartilage. Although controlling overexpression of the transfected gene remains a problem, this is a promising complementary modality for engineering cartilage. VIII DISCUSSION The primary goal of engineering cartilage as a therapeutic approach is to restore the physiological conditions of an affected or defective tissue in the body. Cartilage tissue is distributed widely in the human body and possesses an organization related to the specific demand of a particular anatomical region. In selecting the proper material for engineering cartilage, the functional demands of the replacement tissue must be considered. In summary, there is a multitude of scaffolds, both naturally occurring and synthetic, that are suitable for engineering cartilage. Investigators have shown that the characteristics of the neocartilage differ significantly depending upon which scaffold is employed. There are also large differences when a single scaffold is tested in vitro as opposed to in vivo. Moreover, the addition of other materials internally or externally to the cartilage composite influences the physical and biomechanical properties of the newly formed tissue. The results achieved so far are extremely encouraging and motivate further investigate efforts in the field. The biomechanical composition and, more importantly, the biomechanical properties of the native tissue still represent the ideal replacement tissue. To achieve this goal, however, further studies are needed to duplicate cartilage tissue in the way Mother Nature has perfectly designed. REFERENCES 1. Public Health Service. 1990. Healthy People 2000. U.S. Department of Health and Human Services. 2. Moss M. L., Moss-Salentijn L. 1983. Vertebrate cartilages. In: Cartilage: Structure, Function and Biochemistry, Hall B. K., Ed., Academic Press: New York, pp. 1–30. 3. Amstrong C. G., Mow V. C. 1980. Friction, lubrication and wear of synovial joints. In: Scientific Foundations of Orthopaedics and Traumatology, Owen R., Goodfellow J., Bullough P., Eds. Heinemann: London, pp. 223–232. 4. Happey F. 1980. Studies of the structure of the human intervertebral disc in relation to its functional and aging processes. In: The Joints and Synovial Fluids, Sokoloff L., Ed. Vol. 2, Academic Press: New York, pp. 95–137. 5. Naylor A. 1980. The design and function of the human intervertebral discs. In: Scientific Foundations of Orthopaedics and Traumatology, Owen R. Goodfellow J., Bullough P. Eds. Heinemann: London, pp. 97–105. 6. Myers E. R., Mow V. C. 1983. Biomechanics of cartilage and its response to biomechanical stimuli. In: Cartilage: Structure, Function and Biochemistry, Hall B. K. Ed. Academic Press: New York, pp. 313–341.

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70. Paige K. T., Cima L. G., Yaremchuk M. J., Vacanti J. P., Vacanti C. A. 1995. Injectable cartilage. Plast. Reconstr. Surg. 96:1390–1398. 71. Paige K. T., Cima L. G., Yaremchuk M. J., Vacanti J. P., Vacanti C. A. 1995. Injectable cartilage. Plast. Reconstr. Surg. 96:1390–1398. 72. Sechriest V. F., Miao Y. J., Niyibizi C., Westerhausen-Larsen A., Matthew H. W., Evans C. H., Fu F. H., Suh J. K. 2000. GAG-augmented polysaccharide hydrogel: a novel biocompatible and biodegradable material to support chondrogenesis. J. Biomed. Mater. Res. 49(4):534–541. 73. Lahiji A., Sohrabi A., Hungerford D. S., Frondoza C. G. 2000. Chitosan supports the expression of extracellular matrix proteins in human osteoblasts and chondrocytes. J. Biomed. Mater. Res. 51(4):586–595. 74. Itay S., Abramovici A., Nevo Z. 1987. Use of cultured embryonal chick epiphyseal chondrocytes as grafts for defects in chick articular cartilage. Clin. Orthop. 220:284–303. 75. Smidsrod O., Skjak-Braek G. 1990. Alginate as immobilization matrix for cells. Trends Biotechnol. Mar. 8(3):71–78. 76. Guo J. F., Jourdian G. W., MacCallum D. K. 1989. Culture and growth characteristics of chondrocytes encapsulated in alginate beads. Connect. Tissue Res. 19(2–4):277–297. 77. Ramdi H., Tahri Jouti M. A., Lievremont M. 1993. Immobilized articular chondrocytes: in vitro production of extracellular matrix compounds. Biomater. Artif. Cells Immobilization Biotechnol. 21(3):335–341. 78. Vandenbossche G. M., Bracke M. E., Cuvelier C. A., Bortier H. E., Mareel M. M., Remon J. P. 1993. Host reaction against empty alginate–polylysine microcapsules. Influence of preparation procedure. J. Pharm. Pharmacol. 45(2):115–120. 79. Sims D. D., Butler P. E., Cao Y. L., Casanova R., Randolph M. A., Black A., Vacanti C. A., Yaremchuk M. J. 1998. Tissue engineered neocartilage using plasma derived polymer substrates and chondrocytes. Plast. Reconstr. Surg. 101:1580–1585. 80. Silverman R. P., Passaretti D., Huang W., Randolph M. A., Yaremchuk M. J. 1999. Injectable tissue-engineered cartilage using a fibrin glue polymer. Plast. Reconstr. Surg. 103(7):1809– 1818. 81. Homminga G. N., Bruna P., Koot H. W., van der Kraan P. M., van der Berg W. B. 1993. Chondrocyte behavior in fibrin glue in vitro. Acta Orthop. Scand. 64:441–445. 82. Fortier L. A., Lust G., Mohammed H. O., Nixon A. J. 1999. Coordinate upregulation of cartilage matrix synthesis in fibrin culture supplemented with exogenous insulin-like growth factor-I. J. Orthop. Res. 17:467–474. 83. Silverman R. P., Bonassar L. J., Passaretti D., Randolph M. A., Yaremchuk M. J. 2000. Adhesion of tissue engineered cartilage to native cartilage. Plast. Reconstr. Surg. 105(4):1393– 1398. 84. Park M. S., Cha C. I. 1993. Biochemical aspects of autologous fibrin glue derived from ammonium sulfate precipitation. Laryngoscope 103(2):193–196. 85. Peretti G. M., Randolph M. A., Villa M. T., Buragas M. S., Yaremchuk M. J. 2000. Cell-based tissue engineered allogeneic implant for cartilage repair. Tissue Eng. 6(5):567–576. 86. Peretti G. M., Randolph M. A., Zaporojan V., Bonassar L. J., Xu J. W., Fellers J., Yaremchuk M. J. 2001. A biomechanical analysis of an engineered cell–scaffold implant for cartilage repair. Ann. Plast. Surg. 46(5):533–37. 87. Peretti G. M., Bonassar L. J., Zaporojan V., Xu J. W., Randolph M. A., Yaremchuk M. J. 2000. A cell-based implant for cartilage repair: a biomechanical analysis. Abstract of the Third Biennial Meeting of the Tissue Engineering Society. Tissue Eng. 6(6):699. 88. Xu J. W., Peretti G. M., Nazzal J. A., Kirchhoff C. H., Randolph M. A., Yaremchuk M. J. 2000. Tissue engineered cartilaginous composite using autologous vital cartilage chips and fibrin glue ploymer. Abstract of the Third Biennial Meeting of the Tissue Engineering Society. Tissue Eng. 6(6):687.

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89. Sims C. D., Butler P. E., Casanova R., Lee B. T., Randolph M. A., Lee W. P., Vacanti C. A., Yaremchuk M. J. 1996. Injectable cartilage using polyethylene oxide polymer substrates. Plast. Reconstr. Surg. 98(5):843–850. 90. Elisseeff J., Anseth K., Sims D., McIntosh W., Randolph M., Yaremchuk M., Langer R. 1999. Transdermal photopolymerization of poly(ethylene oxide)–based injectable hydrogels for tissue-engineered cartilage. Plast. Reconstr. Surg. 104(4):1014–1022. 91. Harris J. M. 1992. Poly(ethylene glycol) chemistry. In: Topics in Applied Chemistry, Katritzky A. R., Sabongi G. J., Eds. Plenum Press: New York, p. 385. 92. Sawhney A. S., Pathak C., Hubbell J. A. 1993. Bioerodible hydrogels based on photopolymerized poly(ethylene glyco)-co-poly(-hydroxy acid)diacrylate monomers. Macromolecules 26:581. 93. Hubbell J. A., Massia S. P., Desai N. P., Drumheller P. D. 1991. Endothelial cell–selective materials for tissue engineering in the vascular graft via a new receptor. Biotechnology (NY) 9(6):568–572. 94. Hill-West J. L., Chowdhury S. M., Slepian M. J., Hubbell J. A. 1994. Inhibition of thrombosis and intimal thickening by in situ photopolymerization of thin hydrogel barriers. Proc. Natl. Acad. Sci. USA 91(13):5967–5971. 95. Hill-West J. L., Chowdhury S. M., Sawhney A. S., Pathak C. P., Dunn R. C., Hubbell J. A., 1994. Prevention of postoperative adhesions in the rat by in situ photopolymerization of bioresorbable hydrogel barriers. Obstet. Gynecol. 83(1):59–64. 96. Elisseeff J., Anseth K., Sims D., McIntosh W., Randolph M., Langer R. 1999. Transdermal photopolymerization for minimally invasive implantation. Proc. Natl. Acad. Sci. 96:3104– 3107. 97. Amiji M., Park K. 1992. Prevention of protein adsorption and platelet adhesion on surfaces by PEO/PPO/PEO triblock copolymers. Biomaterials 13(10):682–692. 98. Ashiku S., Randolph M. A., Vacanti C. A., Mathisen D., Yaremchuk M. A. 1997. European Tissue Repair Society, 2nd ETRS Consensus Meeting, August 20–22, Frieburg, Germany. 99. Arevalo-Silva C. A., Eavey R. D., Cao Y., Vacanti M., Weng Y., Vacanti C. A. 2000. Internal support of tissue-engineered cartilage. Arch. Otolaryngol. Head Neck Surg. 126:1448–1452. 100. Nishibe T., Yasuda K., Ohkashiwa H., Watanabe S., Okuda Y., Tanabe T., 2000. High-porosity expanded polytetrafluoroethylene grafts for thoracic vena cava replacement with or without an omentum wrap. Surg. Today 30:631–635. 101. Giovanni A., Vallicioni J. M., Gras R., Zanaret M. 1999. Clinical experience with Gore-Tex for vocal fold medialization. Laryngoscope 109:284–288. 102. Llado A., Sologaistua E., Guimera J., Marin M. 1999. Expanded polytetrafluoroethylene membrane for the prevention of peridural fibrosis after spinal surgery: a clinical study. Eur. Spine J. 8:144–150. 103. Stanec S., Stanec Z. 1998. Ulnar nerve reconstruction with an expanded polytetrafluoroetylene conduit. Br. J. Plast. Surg. 51:637–639. 104. Ballesta-Lopez C., Bastida-Vila X., Catarci M., Bettonica-Larranaga C., Zaraca F. 1998. Laparoscopic gastric banding for morbid obesity with expanded PTFE: technique and early results in the first 100 consecutive cases. Hepatogastroenterology 45:2447–2452. 105. Phillips E., Dardano A. N., Saxe A. 1997. Laparoscopic repair of abdominal hernias using an ePTFE patch—a modification of a previously described technique. J. Soc. Laparoendosc. Surg. 1:277–279. 106. Maas C. S., Gnepp D. R., Bumpous J. 1993. Expanded polytetrafluoroethylene (Gore-Tex soft-tissue patch) in facial augmentation. Arch. Otolaryngol. Head Neck Surg. 119:1008– 1014. 107. Catanese J., Cooke D., Maas C., Pruitt L. 1999. Mechanical properties of medical grade expanded polytetrafluoroethylene: the effects of internodal distance, density, and displacement rate. J. Biomed. Mater. Res. 48:187–192.

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12 Biodegradable Scaffolds for Meniscus Tissue Engineering Mark A. Sweigart and Kyriacos A. Athanasiou Rice University, Houston, Texas

I INTRODUCTION The meniscus, fibrocartilaginous tissue found within the knee joint, is responsible for shock absorption, load transmission, and stability within the knee joint [1–5]. According to the National Center for Health Statistics, over 600,000 surgeries each year are the result of complications with the meniscus [6]. The meniscus has the intrinsic ability to heal itself; unfortunately, this property is limited to the vascular portions of the tissue [7]. For damage outside of these areas and overall degeneration of the tissue, methods need to be developed that will assist the meniscus in healing itself; tissue engineering is a potential solution. Figure 1 depicts our overall philosophy for addressing the problem of successful regeneration of the knee meniscus. Of particular importance is the choice of the scaffold material. This scaffold must be biodegradable, support tissue growth, and protect the articular cartilage surfaces until new tissue is fully developed. II MENISCUS ANATOMY The meniscus is a tissue consisting of two wedge-shaped semilunar sections of fibrocartilaginous tissue between the tibial and femoral bearing surfaces of the knee joint (Fig. 2). On gross inspection the meniscus is a white, glossy, and smooth tissue; this smoothness is also present at the microscopic level [8]. The peripheral portion of the meniscus is vascularized, whereas the inner portion is avascular [9]. The meniscus is attached at the medial collateral ligament, the meniscofemoral ligaments, the transverse ligament, and the anterior and posterior horns. The anterior and posterior horns are where the meniscus joins with the tibial plate; these attachments are usually considered the most important [10]. These 225

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Figure 1 General principle for tissue engineering the meniscus.

Figure 2 Human meniscus (shown attached to tibia).

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horns are also the site of the highest amount of innervation within the meniscus [11,12]. The rest of the innervation occurs in the outer third of the meniscus [11,12]. A Cells in the Meniscus There are two main zones that the meniscus can be divided into, the superficial zone and the deep zone. The cell type present in each of these zones is different. The superficial zone contains cells that are oval or fusiform, have few processes, and feature a scant amount of cytoplasm, resulting in the nucleus of the cell looking abnormally large [8,9]. The deep zone cells are rounded or polygonal are usually alone, but occasionally groups of two or three can be found, and have a large amount of rough endoplasmic reticulum [8,9]. In 1985 Webber and coworkers [13] coined the term fibrochondrocytes to describe the complex fibroblastic and chondrocytic nature of these cells. B Extracellular Matrix of the Meniscus The extracellular matrix can be separated into four different categories: water, fibrillar components, proteoglycans, and adhesion glycoproteins. Human meniscal tissue has been shown to be 72% water, 22% collagen, 0.8% glycosaminoglycans, and the rest made up of DNA and adhesion molecules [14]. These numbers can vary depending on age, animal, and location within the tissue [15,16]. 1 Fibrillar Components The main type of fibrillar component found in the meniscus is collagen. Collagen types I, II, III, V, and VI have been found within meniscal tissue and account for 60–70% of the dry weight [15]. Type I collagen is by far the most predominant, accounting for 90% of the collagen within the tissue [17]. A study done by Cheung [18] has shown that in bovine menisci the outer two-thirds of the menisci is predominantly type I collagen, whereas the inner one-third is 60% type II collagen and 40% type I collagen. The meniscus has a unique collagen structure orientation that is related to its function and consists of three different layers. The superficial layer consists of a thin layer of randomly oriented fibers [19]. The lamellar layer, situated just inside the superficial layer, also consists of randomly oriented fibers, with the exception of the peripheral portions at the anterior and posterior sections; here the fibers are oriented radially [19]. The deep zone consists of circumferentially oriented fibers with a small amount of radially oriented fibers, also referred to as tie fibers [19]. The unique fiber orientation found in the knee meniscus leads to the tissue’s exceptional properties, which will be described later. 2 Proteoglycans Proteoglycans are responsible for hydration within the meniscus and the compressive properties of the tissue [20,21]. The concentration of proteoglycans in meniscal tissue is eightfold less than the concentration found in articular cartilage [22,23]. Proteoglycans are responsible for tissue hydration and various studies have been performed on this component [21]. The inner two-third of the meniscus produces more proteoglycans than the outer third, and the lateral side produces more proteoglycans than the medial side, though the glycosaminoglycan makeup of the proteoglycans stays the same at all of these locations [24,25]. Normal human meniscal tissue consists of 40% chondroitin-4-sulfate, 10–20% chondroitin-6-sulfate, 20–30% dermatan sulfate, and 15% keratan sulfate [26].

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3 Adhesion Glycoproteins As the name suggests, adhesion molecules are partly responsible for binding with other matrix molecules and cells. There are three of these molecules that have been identified within the meniscus: type VI collagen, fibronectin, and thrombospondin [15]. While the exact nature of these three glycoproteins has not yet been discovered, the RGD sequence, which plays a central role in cell adhesion, has been found in type VI collagen, fibronectin, and thrombospondin [15]. III MENISCAL BIOMECHANICS Functionally, the meniscus is a shock absorber, helps with load bearing and transmission in the knee joint, improves stability in the knee, and helps with lubrication [1–5]. Because of all these different functions and the geometry of the tissue, the meniscus is subjected to compressive, tensile, and shear stress. Whenever a load is applied to the knee joint, the meniscus is compressed, but due to its wedge-shaped architecture it is also displaced away from the center of the femoral condyles, resulting in tensile stress because of the anterior and posterior attachment to the tibial plate [4,27,28]. If these attachments are not present, then the biomechanical function of the meniscus is altered and degeneration of the tissues in the area can result [29–31]. In terms of shear properties of the meniscus, it is known that they depend heavily on the collagen orientation of the meniscus and the low circumferential shear strength is thought to be partly responsible for the occurrence of longitudinal tears [32,33]. A Movement of the Meniscus and Force Transmission The menisci partially cover the articular cartilage on the tibial plate and are responsible for absorbing some of the load transmitted through the knee. Application of these loads causes the meniscus to be displaced, and different flexion angles also cause displacement of the tissue [3,4,27,28]. In general, the lateral meniscus is displaced more than the medial meniscus during compression, showing that while the menisci both have the same general function they do react differently [27]. B Mechanical Properties of Meniscal Tissue Tensile, compressive, and shear tests have been run on meniscal tissue. Tensile tests run by Proctor and coworkers [16] have shown the following: 1. Tensile specimens taken from the surface layer, where collagen orientation is random, are isotropic. 2. Tensile specimens taken from the deep zone, where the majority of the collagen fibers are oriented circumferentially, are anisotropic and are the stiffest in the circumferential direction. 3. Circumferential specimens from the posterior section are stiffer than specimens from the anterior section. Tensile tests of the attachments have also been performed, showing that on a rabbit model the anterior attachment is stronger than the posterior attachment and both of these attachments are vital to keep mechanical function [31,34]. Compression tests of the meniscus, although not performed as frequently as tensile tests, have yielded a wealth of information.

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It has been shown that the meniscus has a lower compressive stiffness and lower permeability than articular cartilage, suggesting that the meniscus is an excellent shock absorber [35]. Joshi and coworkers [36] have shown a wide difference in the compressive characteristics of the meniscus between different animals, suggesting caution when designing a scaffold for a particular animal model. The shear properties of the meniscus are theorized to contribute to the larger occurrence of longitudinal tears [35,37]. It has been shown that shear properties are anisotropic, with the stiffness being the lowest in the planes that run parallel to the major collagen orientation [35]. These complex mechanical properties of the meniscus need to be taken into consideration when designing a scaffold that will emulate the native tissue. IV INJURIES AND REPAIR OF THE MENISCUS Injuries to the meniscus usually consist of tears in the tissue, though separation from the tibial attachments and degeneration of the tissue also occur [5,7,31,38,39]. Different styles of tears can occur, such as longitudinal and bucket-handle tears (most common), radial tears, and complex tears [38,39]. In 1936, King [7] was the first to show that tears in the vascularized portion of the meniscus would naturally heal, whereas tears in the avascular portion would not. It has also been shown that longitudinal tears, if they heal, restore native mechanical function, but radial tears, where the collagen structure is disrupted, do not have restored mechanical function after healing [5]. While some repair techniques, such as suture, meniscal arrows, fibrin sealant, laser welding, and abrasion therapy, can help heal longitudinal tears in the vascular region, other techniques need to be developed to heal avascular tears and other complications [40,41]. Tissue engineering offers one way to do this. V TISSUE ENGINEERING OF THE MENISCUS When compared to other musculoskeletal tissues such as bone or articular cartilage, there is a dearth in the number of studies done to date in the attempt to tissue engineer the meniscus. Many components, such as cells, growth factors, animal models, culturing conditions, and evaluation techniques, must be considered when attempting to tissue engineer the meniscus. Of particular importance is the scaffold, on which more information on past and current efforts is given here. A Biodegradable Scaffolds Several different biodegradable scaffolds have been used to tissue engineer the meniscus. Results have varied, with some materials warranting no further research to other materials that are still undergoing testing in animal models and in humans. All of the scaffolds that have been attempted can be separated into two major categories, natural and synthetic. Table 1 gives a summary of different scaffold materials that have been attempted and their effectiveness in tissue engineering the meniscus. 1 Natural Scaffolds Four natural scaffolds have been attempted for tissue engineering of the meniscus: collagen, small intestine submucosa (SIS), periosteal tissue, and perichondrial tissue. Of these four, periosteal tissue had the least acceptable results. In a study done by Walsh and

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Material and reference

Seeded?

Meniscal Scaffolds Test duration

Animal model

Degree of success

Natural scaffolds Perichondral Tissue (1998) [43] Periosteal Tissue (1999) [42] SIS (1999) [44] Collagen (I) sponge (1999) [42] Collagen (I) sponge (1999) [42]

No

12 month

Ovine

Low tensile modulus

No

24 week

Rabbit

No No

12 weeks 24 weeks

Canine Rabbit

Yes (MSCs)

24 weeks

Rabbit

Collagen (I)–GAG (1999) [45]

Yes (meniscal)

3 weeks

In vitro

Collagen (II)–GAG (1999) [45] Collagen (I)–GAG (1992, 1997, 1999) [47,49,50]

Yes (meniscal) No

3 weeks

In vitro

3 years

Canine, human

Bone present in repair tissue Promising results Fibrous tissue, osteoarthritic degeneration Fibrocartilaginous tissue, osteoarthritic degeneration Cells at periphery of scaffold, contraction of scaffold Cells throughout scaffold, good GAG content Phase II clinical trial, slight shrinkage

Synthetic Scaffolds PU (1996) [54] PU (1996) [53] PU (1996) [55] PLLA--capralactone (1997) [56] PGA (1997) [57] PLGA (2000) [58]

No No No No

28 weeks 50 weeks 52 weeks 26 weeks

Canine Canine Canine Canine

Degeneration of AC Degeneration of AC Degeneration of AC Ingrowth of fibrocartilage

Yes (meniscal) Yes (meniscal)

16 weeks Subcutaneous 6 weeks in situ

Rat

Resembled meniscal tissue, lower compressive modulus Organized collagen matrix, good proteoglycan amount

Ovine

coworkers [42] a partial anterior medial meniscectomy was performed in a rabbit model and the defect was filled with a periosteal autograft. After unrestricted activity, the rabbits were sacrificed at 6, 12, and 24 weeks to evaluate the regenerated tissue via gross and histological observations [42]. Results at 6 weeks were promising, but at 12 and 24 weeks bone was noted in the repair tissue and there were severe degenerative changes in the joint [42]. The results from a perichondrial tissue study were not much better [43]. In this study a complete medial meniscectomy was performed on an ovine model and then the defect was filled with autologous perichondrial tissue obtained from the lower rib of the animals [43]. Sacrifice was performed at 3, 6, and 12 months; the regenerated tissue was examined by gross examination, histology, microscopy, and biomechanical testing [43]. The regenerated meniscus showed good results grossly and microscopically, though the histology showed a small area of calcification in the middle of the regenerated tissue [43]. Of greatest concern was the mechanical properties of the regenerated tissue, because the failure stress and tensile modulus of the regenerated tissue were significantly lower than native meniscal tissue

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[43]. Small intestine submucosa is another material that has undergone brief testing as a scaffold for tissue engineering the meniscus [44]. Cook and coworkers [44] have performed a study where porcine SIS was used to fill the defect created by a partial medial meniscectomy in a canine model. The study, which was performed over a 12-week period, showed similarity to native tissue with respect to type II collagen content, zonal architecture, and overall size [44]. No mechanical testing of the tissue was performed and, to the authors’ knowledge, no studies over a longer term have been performed [44]. By far the most promising results for a natural scaffold have come from using collagen as a base material for the scaffold. Walsh and coworkers [42] have performed a study where the effect of a bovine type I collagen sponge was tested to evaluate the regenerative effects under the same conditions listed previously for the periosteal autograft [42]. They found that the sponge gave better results than the periosteal tissue, though the repair tissue was largely fibrous [42]. They also found that better results could be obtained by seeding the scaffolds with mesenchymal stem cells, though this is the only study that had mesenchymal stem cells seeded on a natural scaffold [42]. One study, done by Mueller and coworkers [45], consisted of seeding meniscal cells on type I collagen–GAG scaffolds and type II collagen– GAG scaffolds and incubating them over a 21-day period. The scaffolds were evaluated via histology, immunohistochemistry, GAG and DNA analysis, and the degree of matrix contraction [45]. Results showed that the type I collagen–GAG scaffold exhibited cells growing at the boundaries of the scaffold only, whereas the type II collagen–GAG scaffold exhibited cells throughout [45]. The type II matrix showed increased GAG and DNA content when compared to the type I matrix at the end of the 3-week trial [45]. The type II matrix also did not show evidence of significant shrinkage, whereas the type I matrix shrunk to half its initial size [45]. Walsh and coworkers [45] concluded that the type II matrix should be tested in vivo, though to the authors’ knowledge this has never been done. The best results appear to have come from using a collagen–GAG scaffold developed by Stone, Rodkey, and coworkers [6,46–50]. To create this scaffold, collagen is harvested from bovine Achilles tendons, purified, and then molded into a circumferential orientation via manual mold rotation [46,47]. After a series of in vitro studies, the scaffold was tested in both the porcine and canine animal model [6,47,48]. The initial study, using an immature porcine model, showed that the scaffold had no ill effects toward the healing of the meniscus [48]. Following this study, a mature canine model was tested over a 12-month period, where 80% of the meniscus was removed and replaced with the scaffold. Histological and biochemical tests showed that the scaffold induces regeneration of the meniscus [47]. The excellent results of these studies has led to phase II clinical trials of this product [50]. The phase I and II clinical trials consisted of removing the damaged portion of the meniscus, trimming the scaffold to the appropriate size, and then suturing it in place, with the replaced section varying between 35–85% of the total meniscus [49,50]. Nine patients finished the phase I 36-month trial, and results showed a decrease in the amount of pain noted by the patient and an increase in mobility [49]. The phase II clinical trial, which consisted of eight patients, lasted for at least 24 months and showed similar trends to the phase I trial [50]. Gross examination by arthroscopy showed tissue regeneration in the patients and preservation of the joint articulating surfaces [50]. No adverse effects were noted with the use of this collagen scaffold [50]. 2 Synthetic Scaffolds There is a wider variety of synthetic scaffolds that have been tried, though none of them has yet reached the state that collagen scaffolds have in the development cycle. A study done

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by Klompmaker and coworkers [51] showed that the optimal pore size in a scaffold for inducing ingrowth of fibrocartilage is 150–500 m. This group has also done a large amount of research on using an aliphatic polyurethane for a biodegradable synthetic scaffold [52–54]. This material, which degrades 50–75% after 1 year, was first used in the attempt to improve healing in an avascular longitudinal tear of a canine’s meniscus [52]. The porous polyurethane (86% porosity, 31% being macropores from 250–300 m and the rest being micropores under 90 m) was sutured into the defect and healing was evaluated over a 52-week period [52]. Histological and biochemical tests showed both type I and type II collagen present, and some healing of the tear did occur [52]. Degeneration of the articular cartilage surface was also noted in some of the animals [52]. The porous polyurethane was later used to replace the lateral canine meniscus, though there were a larger number of macropores (43%) in this study, and the compression modulus was tested and found to be 150 kPa [54]. The scaffold was implanted and the animals were unrestricted for the 28week period of the study [54]. Histology showed the presence of fibrocartilage in the scaffold after 28 weeks, though degeneration of the articular cartilage–bearing surfaces did occur [54]. The investigators theorized that this degeneration could be due to the long period of time that elapses before there is sufficient tissue ingrowth into the scaffold [54]. This same group did another lesion study very similar to the first study where they found that vascularization approaches a lesion, and then when the wound is healed the vascularization retracts [53]. Unfortunately, the aromatic polyurethane used in the previous studies releases toxic particles during degradation; therefore the material was altered and an aliphatic polyurethane was used [55]. The aliphatic polyurethane scaffold had a compression modulus of 150 kPa and a porosity of 65%, with 35% of the pores being between 250–300 m and 35% of the pores ranging in size from 50–90 m [55,56]. Results over a 52-week study showed that fibrocartilage was present by 18 weeks after implantation in the canine model, though the tear strength of the implant was weak, leading to loosening of the implant [55]. Complex suturing techniques were used to help alleviate this problem [55]. Unfortunately, at the end of the study some degeneration of the articular cartilage was noted, and they theorized that the scaffold needed a higher compressive modulus to help alleviate this problem [55]. The next study done by this group used a 50:50 copoly(L-lactide/-capralactone) scaffold with varying compressive modulus (40 and 100 kPa) [56]. This material was sutured into a partial defect in the lateral meniscus of the canine and tissue ingrowth was determined; an aliphatic polyurethane, like the one used in the cited study, was also used as a comparison [56]. The copolymer with a compression modulus of 40 kPa had no fibrous tissue ingrowth, whereas the 100-kPa sample had a fibrous tissue ingrowth of 50–70%, and the aliphatic polyurethane had 80–100% fibrous tissue ingrowth [56]. Adhesion of the implant to the native tissue was noted to be higher in the 50:50 co-poly(L-lactide/-capralactone) implants and was theorized to be due to its faster degradation time when compared to the aliphatic polyurethane [56]. Ibarra and coworkers [57,58] have done a series of studies where they use seeded polyglycolic acid (PGA) or poly(lactic-co-glycolic) acid (PLGA) scaffolds in the attempt to tissue engineer the meniscus. Due to the poor mechanical properties of these materials they seed the scaffolds for a few days, implant them subcutaneously for a period of time (depending on the animal model), and then place them in the defect. In one of their preliminary studies, bovine fibrochondrocytes were harvested and expanded in a supplemented Ham’s F-12 culture media and then seeded onto PGA scaffolds at a concentration of 2.5  107 cells/mL [57]. These scaffolds were then implanted subcutaneously in nude mice for a period of 16 weeks [57]. At the end of 16 weeks the tissue from the implant was

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found to grossly and histologically resemble meniscal tissue [57]. Biomechanical tests of some of these subcutaneous implants have also been performed; results showed a compressive modulus that was about 40% of the native tissue value [58]. The most recent test performed by the group consisted of implanting a seeded PGLA scaffold in a sheep model for 4 weeks subcutaneously and 6 weeks in situ [58]. The resultant tissue was rich with proteoglycans and had a organized collagen fiber matrix [58]. These results are quite promising, though more studies need to be performed. In one study, done by Webber and coworkers [59], it was shown that the RGD peptide enhances the attachment of canine fibrochondrocytes to artificial surfaces. This shows great potential for enhancement of current scaffolds that are under development. VI CONCLUSION Successfully tissue engineering the meniscus would be a great help for the treatment of meniscal defects, but much needs to be done before this state can be reached. In particular, scaffolds need to undergo more testing, and the repair tissue that these scaffolds creates needs to be tested grossly, histologically, biochemically, and biomechanically. If none of the current scaffolds being researched give acceptable results, then new meniscus-specific scaffolds need to be developed. Peptides could be used in a scaffold for directed cell attachment, and growth factors could be used for aiding in differentiation and synthesis. The scaffold could also be combined with cells and the tissue could be grown in a bioreactor (Fig. 1). Central to this approach is the use of a suitable scaffold that can allow the implementation of innovative strategies to tissue engineer the meniscus. REFERENCES 1. Hoshino, A., Wallace, W. A. 1987. Impact-absorbing properties of the human knee. J. Bone Joint Surg. Br. 69:807–811. 2. Radir E. L., de Lamotte F., Maquet P. 1984. Role of the menisci in the distribution of stress in the knee. Clin. Orthop. 290–294. 3. Ahmed A. M. 1992. The load-bearing role of the knee meniscus. In: Knee Meniscus: Basic and Clinical Foundations. Mow V. C., Arnoczky S. P., Jackson D. W. Eds. Raven Press: New York, pp. 59–73. 4. Walker P. S., Erkman M. J. 1975. The role of the menisci in force transmission across the knee. Clin. Orthop. 109:184–192. 5. Newman A. P., Anderson D. R., Daniels A. U., Dales M. C. 1989. Mechanics of the healed meniscus in a canine model. Am. J. Sports Med. 17:164–175. 6. Rodkey W. G., Stone K. R., Steadman J. R. 1992. Prosthetic meniscal replacement. In: Biology and Biomechanics of the Trumatized Synovial Joint: The Knee as a Model. Finerman G. A. M., Noyes F. R., Eds. American Academy of Orthopaedic Surgeons: Rosemont, pp. 221–231. 7. King D. 1936. The function of semilunar cartilages. J. Bone Joint Surg. 18:1069–1076. 8. Ghadially F. N., Thomas I., Yong N., Lalonde J. M. 1978. Ultrastructure of rabbit semilunar cartilages. J. Anat. 125:499–517. 9. Ghadially F. N., Lalonde J. M., Wedge J. H. 1983. Ultrastructure of normal and torn menisci of the human knee joint. J. Anat. 136:773–791. 10. Gao J. 2000. Immunolocalization of types I, II, and X collagen in the tibial insertion sites of the medial meniscus. Knee Surg. Sports Traumatol. Arthrosc. 8:61–65. 11. Mine T., Kimura M., Sakka A., Kawai S. 2000. Innervation of nociceptors in the menisci of the knee joint: an immunohistochemical study. Arch. Orthop. Trauma Surg. 120:201–204. 12. Zimny M. L. 1988. Mechanoreceptors in articular tissues. Am. J. Anat. 182:16–32.

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13. Webber R. J., Harris M. G., Hough Jr. A. J. 1985. Cell culture of rabbit meniscal fibrochondrocytes: proliferative and synthetic response to growth factors and ascorbate. J. Orthop. Res. 3:36–42. 14. Herwig J., Egner E., Buddecke E. 1984. Chemical changes of human knee joint menisci in various stages of degeneration. Ann. Rheum. Dis. 43:635–640. 15. McDevitt C. A., Webber R. J. 1990. The ultrastructure and biochemistry of meniscal cartilage. Clin. Orthop. 8–18. 16. Proctor C. S., Schmidt M. B., Whipple R. R., Kelly M. A., Mow V. C. 1989. Material properties of the normal medial bovine meniscus. J. Orthop. Res. 7:771–782. 17. Eyre D. R., Wu J. J. 1983. Collagen of fibrocartilage: a distinctive molecular phenotype in bovine meniscus. FEBS Lett. 158:265–270. 18. Cheung H. S. 1987. Distribution of type I, II, III and V in the pepsin solubilized collagens in bovine menisci. Connect. Tissue Res. 16:343–356. 19. Petersen W., Tillmann B. 1998. Collagenous fibril texture of the human knee joint menisci. Anat. Embryol. (Berlin) 197:317–324. 20. Scott P. G., Nakano T., Dodd C. M. 1997. Isolation and characterization of small proteoglycans from different zones of the porcine knee meniscus. Biochim. Biophys. Acta 1336:254–262. 21. Adams M. E., Hukins D. W. L. 1992. The extracellular matrix of the meniscus. In: Knee Meniscus: Basic and Clinical Foundations. Mow V. C., Arnoczky S. P., Jackson D. W., Eds. Raven Press: New York, pp. 15–28. 22. Adams M. E., Muir H. 1981. The glycosaminoglycans of canine menisci. Biochem. I. 197:385– 389. 23. McNicol D., Roughley P. J. 1980. Extraction and characterization of proteoglycan from human meniscus. Biochem. I. 185:705–713. 24. Collier S., Ghosh P. 1995. Effects of transforming growth factor beta on proteoglycan synthesis by cell and explant cultures derived from the knee joint meniscus. Osteoarthritis Cartilage 3:127–138. 25. Tanaka T., Fujii K., Kumagae Y. 1999. Comparison of biochemical characteristics of cultured fibrochondrocytes isolated from the inner and outer regions of human meniscus. Knee Surg. Sports Traumatol. Arthrosc. 7:75–80. 26. Verbruggen G., Verdonk R., Veys E. M., Van Daele P., De Smet P., Van den Abbeele K., Claus B., Baeten D. 1996. Human meniscal proteoglycan metabolism in long-term tissue culture. Knee Surg. Sports Traumatol. Arthrosc. 4:57–63. 27. Bylski-Austrow D. I., Ciarelli M. J., Kayner D. C., Matthews L. S., Goldstein S. A. 1994. Displacements of the menisci under joint load: an in vitro study in human knees. J. Biomech. 27: 421–431. 28. Fu F. H., Thompson W. O. 1992. Biomechanics and kinematics of meniscus. In: Biology and Biomechanics of the Trumatized Synovial Joint: The Knee as a Model. Finerman G. A. M., Noyes F. R., Eds. American Academy of Orthopaedic Surgeons: Rosemont, pp. 153–183. 29. Gao J., Wei X., Messner K. 1998. Healing of the anterior attachment of the rabbit meniscus to bone. Clin. Orthop. 246–258. 30. Alhalki M. M., Howell S. M., Hull M. L. 1999. How three methods for fixing a medial meniscal autograft affect tibial contact mechanics. Am. J. Sports Med. 27:320–328. 31. Gao J., Messner K. 1996. Natural healing of anterior and posterior attachments of the rabbit meniscus. Clin. Orthop. 276–284. 32. Anderson D. R., Woo S. L., Kwan M. K., Gershuni D. H. 1991. Viscoelastic shear properties of the equine medial meniscus. J. Orthop. Res. 9:550–558. 33. Zhu W., Chern K. Y., Mow V. C. 1994. Anisotropic viscoelastic shear properties of bovine meniscus. Clin. Orthop. 34–45. 34. Goertzen D., Gillquist J., Messner K. 1996. Tensile strength of the tibial meniscal attachments in the rabbit. J. Biomed. Mater. Res. 30:125–128.

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35. Fithian D. C., Kelly M. A., Mow V. C. 1990. Material properties and structure–function relationships in the menisci. Clin. Orthop. 19–31. 36. Joshi M. D., Suh J. K., Marui T., Woo S. L. 1995. Interspecies variation of compressive biomechanical properties of the meniscus. J. Biomed. Mater. Res. 29:823–828. 37. Setton L. A., Guilak F., Hsu E. W., Vail T. P. 1999. Biomechanical factors in tissue engineered meniscal repair. Clin. Orthop. S254–S272. 38. Manco L. G., Kavanaugh J. H., Fay J. J., Bilfield B. S. 1986. Meniscus tears of the knee: prospective evaluation with CT. Radiology 159:147–151. 39. DeHaven, K. E., Arnoczky S. P. 1994. Meniscus repair: basic science, indications for repair, and open repair. Instr. Course Lect. 43:65–76. 40. McAndrews P. T., Arnoczky S. P. 1996. Meniscal repair enhancement techniques. Clin. Sports Med. 15:499–510. 41. DeHaven K. E. 1999. Meniscus repair. Am. J. Sports Med. 27:242–250. 42. Walsh C. J., Goodman D., Caplan A. I., Goldberg V. M. 1999. Meniscus regeneration in a rabbit partial meniscectomy model. Tissue Eng. 5:327–337. 43. Bruns J., Kahrs J., Kampen J., Behrens P., Plitz W. 1998. Autologous perichondral tissue for meniscal replacement. J. Bone Joint Surg. Br. 80:918–923. 44. Cook J. L., Tomlinson J. M., Kreeger C. R., Cook 1999. Induction of meniscal regeneration in dogs using a novel biomaterial. Am. J. Sports Med. 27:658–665. 45. Mueller S. M., Shortkroff S., Schneider T. O., Breinan H. A., Yannas I. V., Spector M. 1999. Meniscus cells seeded in type I and type II collagen–GAG matrices in vitro. Biomaterials 20:701–709. 46. Stone K. R. 1989. Prosthetic meniscus. 4,880,429. 47. Stone K. R., Rodkey W. G., Webber R., McKinney L., Steadman J. R. 1992. Meniscal regeneration with copolymeric collagen scaffolds. In vitro and in vivo studies evaluated clinically, histologically, and biochemically. Am. J. Sports Med. 20:104–111. 48. Stone K. R., Rodkey W. G., Webber R. J., McKinney L., Steadman J. R. 1990. Future directions. Collagen-based prostheses for meniscal regeneration. Clin. Orthop. 129–135. 49. Stone K. R., Steadman J. R., Rodkey W. G., Li S. T. 1997. Regeneration of meniscal cartilage with use of a collagen scaffold. Analysis of preliminary data. J. Bone Joint Surg. Am. 79:1770–1777. 50. Rodkey W. G., Steadman J. R., Li S. T. 1999. A clinical study of collagen meniscus implants to restore the injured meniscus. Clin. Orthop. S281–S292. 51. Klompmaker J., Jansen H. W., Veth R. P., Nielsen H. K., de Groot, J. H., Pennings A. J. 1993. Porous implants for knee joint meniscus reconstruction: a preliminary study on the role of pore sizes in ingrowth and differentiation of fibrocartilage. Clin. Mater. 14:1–11. 52. Klompmaker J., Jansen H. W., Veth R. P., Nielsen H. K., de Groot J. H., Pennings A. J., Kuijer R. 1992. Meniscal repair by fibrocartilage? An experimental study in the dog. J. Orthop. Res. 10:359–370. 53. Klompmaker J., Veth R. P., Jansen H. W., Nielsen H. K., de Groot J. H., Pennings A. J., Kuijer R. 1996. Meniscal repair by fibrocartilage in the dog: characterization of the repair tissue and the role of vascularity. Biomaterials 17:1685–1691. 54. Klompmaker J., Veth R. P., Jansen H. W., Nielsen H. K., de Groot J. H., Pennings A. J. 1996. Meniscal replacement using a porous polymer prosthesis: a preliminary study in the dog. Biomaterials 17:1169–1175. 55. de Groot J. H., de Vrijer R., Pennings A. J., Klompmaker J., Veth R. P., Jansen H. W. 1996. Use of porous polyurethanes for meniscal reconstruction and meniscal prostheses. Biomaterials 17:163–173. 56. de Groot J. H., Zijlstra F. M., Kuipers H. W., Pennings A. J., Klompmaker J., Veth R. P., Jansen H. W. 1997. Meniscal tissue regeneration in porous 50/50 copoly(L-lactide/-caprolactone) implants. Biomaterials 18:613–622.

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13 Type I Collagen–Based Template for Meniscus Regeneration Shu-Tung Li ReGen Biologics, Inc., Franklin Lakes, New Jersey and Collagen Matrix, Inc., Franklin Lakes, New Jersey William G. Rodkey ReGen Biologics, Inc., Franklin Lakes, New Jersey and Steadman-Hawkins Sports Medicine Foundation, Vail, Colorado Debbie Yuen* and Peggy Hansen* ReGen Biologics, Inc., Franklin Lakes, New Jersey J. Richard Steadman Steadman-Hawkins Sports Medicine Foundation, Vail, Colorado

I INTRODUCTION Collagen, a group of proteins of similar structural characteristics, is the most abundant protein in the body. It is the major component of bone, skin, ligament, and tendon. Since collagen, particularly type I collagen, possesses the unique physicochemical, mechanical, and biological properties that are suitable for tissue and organ repair, this protein has been extensively researched in the past two decades as a biomaterial for medical implant development. A summary of the biotechnology of collagen and its medical applications can be found in a recent review [1]. Of particular interest among the various medical uses of collagen is the repair of damaged weight-bearing cartilagenous tissues such as the menisci of the knee joint, employing an engineered collagen matrix template to support and guide the meniscus regeneration. * Current affiliation: Collagen Matrix, Inc., Franklin Lakes, New Jersey

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Menisci are crescent-shaped fibrocartilages that are anatomically situated between the femoral condyles and tibial plateau and provide various biomechanical functions within the knee [2]. Injury to the knee frequently results in a tear of meniscal tissue. Repair of the torn tissue in the peripheral vascular region may be accomplished arthroscopically with sutures [3,4]. However, when the injury is in the inner avascular region or when the meniscus is damaged to the extent that suture repair is not feasible, removal of the damaged tissue is indicated [4,5]. It is known that partial or total removal of a meniscus causes force concentration on the opposing articular cartilages and with time results in degenerative arthritis [6–11]. At the present time, there is no commercial product that can be used for the repair of partial meniscetomized tissue. Our research in applying fibrous type I collagen for various tissue repair has led to the working hypothesis of an engineered extracellular matrix to support and guide meniscus regeneration. We hypothesize that meniscus has the intrinsic ability to regenerate provided that the biological environment is suitable for regeneration. Thus, for an engineered collagen matrix to function as a resorbable meniscus template, one must be certain that the design requirements for the product are met. Of particular importance is the biomechanical property of the matrix template since the template serves the biomechanical function of the meniscus. Thus, the initial biomechanical strength of the engineered template and the subsequent biomechanical properties of the regenerated and remodeled tissue must be adequate for the template to function. In addition, the extracellular matrix must be conductive to cells as well as permeable to nutrients. The biological signals (e.g., growth factors) and the cells may be incorporated into the template to accelerate the overall regeneration and remodeling process and may also provide a more ideal biological environment for cellular infiltration and new matrix synthesis. This paper summarizes the design and development of this collagen-based extracellular matrix template implant. II OVERVIEW OF DESIGN REQUIREMENTS FOR A TYPE I COLLAGEN–BASED MENISCUS TEMPLATE In order to design a resorbable, collagen-based template for supporting and guiding meniscus regeneration, a number of critical design requirements must be considered. The following summarizes the rationale for each of these requirements. A Biocompatibility Biocompatibility of the materials used in the engineering of a resorbable meniscus regeneration template and their degradation products are a prerequisite for resorbable implant development. The engineered resorbable meniscus template consists of primarily type I collagen fibers (97.5%w/w) with the incorporation of small amounts of hyaluronic acid (1.25%w/w) and chondroitin sulfate (1.25%w/w) glycosaminoglycans (GAG) to improve hydrophilic and osmotic properties. Soluble type I collagen is long known to be a poor immunogen [12]. A significant level of antibodies against type I collagen cannot be raised without the use of Freund’s complete adjuvant. It is also known that insoluble collagen is even less immunogenic [13–15]. The meniscus template is engineered from purified insoluble type I collagen fibers that are further crosslinked chemically to increase the stability and reduce the immunogenicity in vivo. Numerous in vitro studies have shown that type I collagen in its various forms maintains cell differentiation and support cell proliferation [16–24]. Thus, many implantable type I collagen–based devices have been developed and successfully marketed over the past two decades. Both the hyaluronic acid and chondroitin sulfate have a long history of successful

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clinical use as implantable devices. Nevertheless, the finished collagen-based meniscus template must be safe for implantation. Therefore, the finished prototypes must be subjected to biocompatibility testing recommended by the FDA. Specifically, cytotoxicity, pyrogenicity, immunogenicity, hemolysis, mutagenicity, and short-term implantation must be evaluated for the resorbable type I collagen–based templates. B Physical Dimension The native crescent-shaped meniscus has a thickness of about 7–8 mm at the periphery and gradually tapers to a thin tip at the inner margin, forming a slightly concave triangle in cross-section. The shape of the meniscus is defined by the space available between the femoral condyles and tibial plateau within the synovial cavity. In designing a meniscus template, the physical dimension of the template defines the boundary of tissue ingrowth and regeneration. The collagen-based template should be a porous, C-shaped, and compliant matrix. The size of the matrix to fit the missing meniscus may be critical from the point of view that a more exact size fit may minimize the erosion or delamination of the matrix from mechanical shear during normal movement of the knee. C Apparent Density The apparent density is defined as the weight of the dry matrix in a unit volume of matrix. Thus, the apparent density is a direct measure of the interstitial space, which is not occupied by the matrix material per se in the dry state. For example, for a collagen matrix of an apparent density 0.2 g/cm3, the interstitial space would be 0.86 cm3 for a 1-cm3 total space occupied by the matrix, taking the density of collagen to be 1.41 g/cm3 [25]. The apparent density is also directly related to the mechanical strength of a matrix. In load-bearing applications, the apparent density must be optimized such that the mechanical strength is not compromised for the intended function of the template implant. The apparent density of the collagen matrix template should be in the range such that the matrix is sufficiently porous for cellular ingrowth yet maintaining the mechanical strength for suturing onto the host for implant attachment without migration. As will be seen in the following sections, the density in the range of 0.18 g/cm3 to 0.22 g/cm3 can provide the optimal pore structure and mechanical strength of the matrix for meniscus regeneration. D Pore Structure The dimension of a mammalian fibrogenic cell body and its processes is on the order of 10 to 50 m, depending on the substrate to which the cells adhere [26]. In order for cells to infiltrate into the interstitial space of a matrix, the majority of the pores must be significantly larger than the dimension of a cell such that both the cell and its cellular processes can easily enter the interstitial space. In a number of studies using collagen-based matrices for tissue regeneration, it has been found that pore size plays an important role in the effectiveness of the collagen matrix to facilitate and support host tissue regeneration [27–29]. It was suggested that pore size in the range of 100 to 400 m was optimal for cellular ingrowth [30]. Similar observations were found to be true for porous metal implants in total hip replacement [31]. The question of interconnecting pores may not be a critical issue as collagenases are synthesized by most cells during wound healing and tissue remodeling [32]. The interporous membranes which existed in the noninterconnecting pores should be digested as part of the wound healing and remodeling processes.

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In the design of a template matrix material for meniscal tissue repair and regeneration, pore size greater than 50 m and preferably in the range from 50 to 400 m should be targeted to provide the optimal environment for cellular ingrowth and tissue regeneration. E Mechanical Properties In designing a resorbable template implant for load-bearing application, not only the initial biomechanical characteristics are important in order for the template to function as a temporary meniscus substitute; the gradual strength reduction of the partially resorbed template must be compensated by the strength increase from the regenerated tissue, such that at any given time point the total mechanical strength of the template–new tissue complex is maintained. In order to accomplish this goal, one must first be certain that the initial mechanical characteristics of the template are adequate to permit surgical implantation. The suture pullout strength is important during surgical placement procedures in order to reduce the incidence of implant failure. In addition, adequate and consistent biomechanical properties such as the compressive stiffness and permeability of the regenerated tissue implant complex must be maintained in order to maintain proper biological function of the regenerated tissue. The suture pullout strength of the hydrated template of greater than 1 kg (about 10 Newtons) was found to be sufficient to implant the template by arthroscopic-assisted surgery in simulated placement procedures in human cadaveric knees (personal communication, Stone, 1991). This strength permits the surgical implantation of the template device by an arthroscopically assisted procedure with minimal risk of implant failure and detachment of the implant from the host tissue. Thus, the collagen template should meet the requirement of suture pullout strength of a hydrated template matrix of 1 kg. The compressive aggregate modulus (Ha) and permeability (k), as defined by Mow [33], define the stiffness and permeability of the template implant. The implant should have adequate compressive aggregate modulus to permit conformity of the implant to the knee joint. The compressive aggregate modulus of the hydrated template should be designed using normal meniscus as a guide. The template should be highly porous for cellular ingrowth and permeable to nutrients and macromolecules (see also sections on pore structure and permeability). F Hydrophilicity Hydration of an implant facilitates nutrient diffusion. The extent of hydration will also provide accurate information on the space available for cellular ingrowth. Collagen is a hydrophilic material, and the crosslinked collagen-based matrix generally behaves like an elastic sponge allowing fluid outflow under stress and fluid inflow when the stress is released, as may be observed with normal flexion of the joint. The elastic property of a hydrated matrix can also provide shock absorption capacity. The meniscus template is comprised of crosslinked, hydrophilic type I collagen– GAG material. The matrix material would be readily hydrated when in contact with water or aqueous fluid and should also facilitate nutrient exchange and diffusion in the joint environment. G Permeability The diffusion of nutrients into the interstitial space of a matrix ensures the survival of the cells and their continued growth and synthesis of tissue-specific extracellular matrix [34].

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As the molecular weight of most nutrients and growth factors is less than 50,000, a probe molecule the size of bovine serum albumin (BSA) (MW 67,000) can be used to test the permeability properties of the matrix to nutrients and growth factors. Ideally, the matrix should be readily permeable to macromolecules the size of bovine serum albumin. H In Vivo Stability The rate of template matrix resorption and the rate of new meniscus tissue regeneration must be balanced so that the adequate mechanical properties are maintained. The rate of in vivo resorption of a collagen-based implant can be controlled by controlling the extent of intermolecular crosslinking. The control of the extent of intermolecular crosslinking can be accomplished by using crosslinking agents that can react with amino, guanidino, carboxyl, and hydroxyl groups of collagen under conditions that do not denature the collagen. Glutaraldehyde, formaldehyde, adipyl chloride, hexamethylene diisocyanate, dye-sensitized photo-oxidation, polyepoxy compounds, and carbodiimides are among the many agents used in crosslinking collagen-based implants in solutions. Crosslinking can also be achieved through vapor phase of a crosslinking agent. The vapor phase crosslinking is usually effective using crosslinking agents of high vapor pressures such as formaldehyde and glutaraldehyde. Additionally, crosslinking formation may be achieved by a thermodynamically unfavorable endothermic condensation (dehydration) reaction by heat and vacuum, referred to in the literature as dehydrothermal crosslinking [35]. The hydrothermal shrinkage temperature of the crosslinked matrix has been used as a guide for assessing the in vivo stability of a collagen implant [36–40]. Temperature of shrinkage of hydrated collagen fibers measures the transition of collagen molecules from triple helix to a random coil conformation. This temperature depends on the extent of intermolecular crosslinks formed by chemical means [41]. Generally, the higher the number of intermolecular crosslinks, the higher will be the hydrothermal shrinkage temperature and longer will be the in vivo stability. Due to the specific three-dimensional packing of collagen molecules and the number of reactive groups available for crosslinking by a specific agent [42–44], there will be a maximum number of intermolecular crosslinks formed under a given experimental condition. Therefore, crosslinking under nonequilibrium conditions may introduce differential degrees of crosslinking, which may result in a differential hydrothermal stability and, thus, a differential in vivo rate of resorption of a given implant. This is particularly manifested in thick and dense implants where the extent of crosslinking is a function of the rate of diffusion of a crosslinking agent [45]. We will target the hydrothermal shrinkage temperature of the template to be in the range of 60 to 70°C. Template having the hydrothermal shrinkage temperature in the range of 60–70°C has an in vivo stability between 6 to 12 months [39,46]. I Sterilization In designing a collagen-based template for load-bearing application, one must carefully choose a method for sterilizing the implant. The current methods for sterilizing collagenbased implants include ethylene oxide, gamma irradiation, and electron beam irradiation. Each method has its advantages and limitations and requires an in-depth evaluation before adapting a particular mode of sterilization. Each of the cited design requirements is evaluated. The interdependency of the design requirements is also considered in an effort to maximize the safety and efficacy of the template.

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III MATERIALS AND METHODS A Type I Collagen Fibers and Glycosaminoglycans 1 Type I Collagen Fibers Type I collagen fibers were isolated from 12- to 18-month-old bovine achilles tendon. The method for preparing the purified type I collagen fibers has been described [39]. The acceptable criteria of purity are based on a comparative study with the commercial collagen implants, which are known to have a long history of successful clinical use. Since the purified collagen fibers have a very low level of noncollagenous impurities and since many current analytical methods have a sensitivity of 3%, the determination of impurities of collagen preparation must be accomplished with a number of analytical methods which simultaneously show a minimal content of impurities. Specifically, amino acid composition, hexosamine content, total neutral sugar content, and hydroxyproline content for the purified collagen were determined. The amino acid composition of the purified collagen was conducted at Health Science Services, Inc. (Riverdale, NJ). The hexosamine content in the purified collagen sample was determined according to the method of Cessi and Piliego [47] using glucosamine (Sigma Chemical, St. Louis, MO) as a standard. The total neutral sugar content was determined based on the anthrone method of Trevelyn and Harrison [48]. Glucose (Sigma Chemical, St. Louis, MO) was used as a standard. Hydroxyproline content was determined according to the method of Bergman and Loxley [49]. 2 Glycosaminoglycans Hyaluronic acid (MW 5  10 5 to 7.5  105) was obtained from LifeCore Biomedical Corp. (Chaska, MN). The hyaluronic acid was of clinical grade and was used without further treatment. Chondroitin sulfate was obtained from Seikagaku Corporation (Tokyo, Japan). The chondroitin sulfate was of medical grade and was also used without further treatment. To further insure the quality of the GAGs, the potential protein contamination in the hyaluronic acid and chondroitin sulfate was determined by a protein microassay using the Coomassie Plus Protein Assay Reagent (Pierce, Rockford, IL), which is based on the Bradford method of protein analysis [50]. Bovine serum albumin (Sigma Chemical, St. Louis, MO) was used for the preparation of the standard solutions. B Engineering of Type I Collagen–Based Meniscus Template The general procedure for engineering a type I collagen–based meniscus template has been described [39]. Briefly, the purified collagen fibers were swollen in the presence of equal quantities of hyaluronic acid and chondroitin sulfate (2.5% GAGs of the dry weight of the finished meniscus implant) at pH 2.5–3.0 and homogenized. The swollen collagen–GAG fibers were coacervated by the addition of 0.6% NH 4OH to the isoelectric point of the purified collagen (pH 5.0). The coacervated fibers were dehydrated, molded, lyophilized, chemically crosslinked, and sterilized. Two crosslinking agents were evaluated during the development phase. Glutaraldehyde, at a concentration of 0.1%, pH 7.4, was used initially and was abandoned later due to nonuniform crosslinking of the matrix. Crosslinking with formaldehyde vapor was then selected as a potential crosslinking method for the dense matrix. The matrix was crosslinked with a formaldehyde vapor (generated under saturated vapor of 2.0% formaldehyde solution at 22°C) for 24 to 48 h at room temperature. The formaldehyde crosslinked matrices

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were then further crosslinked dehydrothermally for 24 h at 100°C under high vacuum (30 to 60 mHg). The residuals of crosslinking agent were removed by extensive rinsing in distilled water to a level which did not show cytotoxicity in vitro according to a U.S. Pharmacopeia (USP) cytotoxicity testing protocol. 1 Method of Sterilization Two sterilization methods were evaluated during the development of the collagen meniscus template. The ethylene oxide sterilization was initially evaluated as a potential method. Standard ethylene oxide sterilization procedure by a contract sterilization facility was followed and residuals of ethylene glycol, ethylene oxide, and ethylene chlorohydrin were tested. The ethylene chlorohydrin (a toxic chemical) was found to be consistently high despite a thorough evacuation. Consequently, it was concluded that ethylene oxide was not a suitable sterilization method. The method of gamma irradiation sterilization was then pursued. It is known that gamma irradiation on a collagenous material induces both the intermolecular crosslinking formation (structure stabilizing effect) and peptide bond scission (structure destabilizing effect) [51,52]. The relative contribution of these effects depends on the microenvironment of the collagen and the dose of the gamma ray to which the collagen is being exposed. Generally, at a dose of 25 kGy of irradiation, the structure destabilizing effect dominates. It is therefore of particular importance to evaluate the effect of gamma irradiation on the key structure parameters of collagen such that the design parameters discussed are not significantly affected. It has been documented that water molecules stabilize the collagen molecules by forming water-bridged hydrogen bonds between the collagen polypeptide chains forming interchain hydrogen bonds [53]. The water molecules can also stabilize the native collagen fibers via intermolecular hydrogen bond formation [43,54,55]. In the evaluation of the effect of gamma irradiation on the collagen template, we designed a study which permitted us to evaluate the structure stabilizing effect of the water molecules on the key structure parameters of the collagen template as a function of the dose of irradiation. The water content of the collagen templates was hydrated to a level from 0 to 200% (weight of water per dry weight of matrix). The dry weight of a matrix (0% water) was operationally defined as the matrix dried under vacuum for 24 h. The dose of irradiation ranges from 3.2 to 26.7 kGy. The key structural parameters evaluated were hydrothermal shrinkage temperature, a parameter which is related to the stability of the triple helix and the intermolecular interactions; swelling properties, which are related to the electrostatic, the intermolecular, and the interfibrillar interactions; mechanical strength, which is related to the intermolecular and interfibrillar interactions and the structural cohesiveness; and trypsin susceptibility, which is related to the stability of the triple helix. The method of characterization for each of the key structural parameters is described in the next section and the preliminary work has been published [56,57,83]. IV BIOCOMPATIBILITY STUDIES A Immunological Studies The potential immunological response to the type I collagen–based template was evaluated in two animal models. The humoral immune response in terms of antibody production

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against the type I collagen–based template was investigated in a rabbit model. The cell-mediated immune response in terms of lymphocyte proliferative activity was investigated in a mouse model. 1 Humoral Immunity The humoral immune response to the type I collagen–based template was evaluated by subcutaneous implantation of the template material as antigen. Positive controls were also developed where the implants were pulverized and injected subcutaneously in the presence of Freund’s complete adjuvant. The antibody production against the implant material was assessed by enzyme-linked immunosorbant assay (ELISA) of serum from the test animals [58,59]. Results were compared with those from preimmune sera drawn prior to immunization. 2 Cell-Mediated Immunity T-lymphocytes proliferate when exposed to an immunogen. The thymus contains noncommitted T-lymphocytes. By comparing the lymphocyte response from the lymph node and spleen lymphocytes (which may be activated by the type I collagen–based template) with the noncommitted lymphocytes in the thymus, the possible long-term immunogenicity of an implant can be gauged [60,61]. B Cytotoxicity Cytotoxicity testing was conducted according to the USP Agarose Overlay protocol. This test was conducted at NAmSA (Northwood, OH). In brief, the type I collagen–based templates (1.6 g) were extracted in 8 mL of 0.9% saline solution at 37°C for 24 h. A monolayer of L-929 mouse fibroblast cells was grown to confluency and overlaid with minimum essential medium (MEM) supplemented with serum, antibiotics, neutral red, and agarose. A 0.1-mL portion of the extract on a filter paper disc was placed on the solidified overlay surface. A filter paper disc saturated with 0.1 mL 0.9% saline was used as a negative control. A 0.5 cm  0.5 cm piece of latex material was used as a positive control. Following incubation for 24 h, the culture was macroscopically examined for evidence of cell decolorization to determine the zone of cell lysis. No change in cell morphology in the proximity to the test material was scored as nontoxic. C Hemolysis The hemolysis test was conduct at NAmSA (Northwood, OH) according to USP. Briefly, a 60.0-cm2 portion of the type I collagen–based template was placed in 20 mL of 0.9% NaCl solution and extracted at 37°C for 24 h. The extract was divided into individual tubes of 5 mL each and allowed to cool to room temperature. A 0.1-mL sample of rabbit blood previously collected in a vacuum tube containing ethylenediamine tetraacetic acid (EDTA) was added to the extract and the positive and negative control tubes. The tubes were inverted gently to mix the contents, then placed in a constant temperature water bath at 37°C for 1 h. The blood saline mixture and the positive and negative controls were then centrifuged for 10 min at a speed of not less than 1000g. The absorbance of each test article solution, positive control (10 mL water and 0.2 mL blood), and negative control (10 mL 0.9% NaCl solution and 0.2 mL blood) were recorded.

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D Pyrogenicity This test was performed by Marypaul Laboratories (Sparta, NJ). The bacterial endotoxins that might be present in the type I collagen–based template were tested using the USP Limulus Amebocyte Lysate (LAL) test method. Determination of the reaction end point was made with dilutions from the sample extract in direct comparison with parallel dilution of a reference endotoxin. Amounts of endotoxin are expressed in defined endotoxin units (EU). E Mutagenicity The Ames mutagenicity test was conducted at NAmSA (Northwood, OH). A Salmonella/ mammalian microsome mutagenicity test was conducted to determine whether a saline extract of the type I collagen–based template would cause mutagenic changes in histidine-dependent mutant strains of Salmonella typhimurium. The method of Ames [62]. was followed with modification for the test of collagen template extracts. V ANIMAL STUDIES A Animal Model We evaluated the design concept in two implantation models in dogs. The first model was a “die punch model,” where two 4-mm-diameter full-thickness discs were removed from the medial meniscus, one from the anterior third and the other from the posterior third, and replaced either with a type I collagen–based template or a fibrin clot or were left unfilled. This model has the advantage that the implants are well protected under a controlled environment. Therefore, a longer-term study can be carried out without the risk of losing specimens due to potential biomechanical failures. Consequently, this model permits evaluation of various safety and efficacy parameters of the template in a controlled and systematic manner. In the die punch model, we evaluated the gross appearance and the biomechanical, biochemical, and histologic characteristics of the regenerated meniscus tissue. The gross appearance of the explant provided information on the stability of the template. The biomechanical characterization of the explant provided information on the biomechanical characteristics of the template during the period of resorption and ingrowth of new tissue. The biochemical characterization of the explant provided information on the nature of the regenerated tissue. The histological characterization of the explant provided the following information: (1) the host tissue response to the template; (2) the extent of cellular ingrowth and template degradation; and (3) the morphological appearance of the cells and ingrown tissue. The second implantation model was a “subtotal resection model,” where 80% of the meniscus was removed leaving a 1 to 2-mm peripheral rim intact for the conduction of cells and nutrients to the matrix template and for suture attachment of the template to the host tissue. The subtotal resection model resembles the human chronic degenerative knees where much of the meniscus has been lost due to trauma and/or degeneration. This study was carried out for 6 months to evaluate the following: (1) biomechanical competence of the template to remain stable in vivo; (2) biomechanical stability of the template in vivo to support and guide tissue ingrowth and for in vivo function; and (3) extent of template re-

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sorption and the cellular ingrowth and tissue maturation. In the subtotal resection model, we evaluated the explants grossly and histologically. Since this surgery is via open arthrotomy, gross examination permitted a thorough evaluation of the healing characteristics, particularly at the posterior horn region where the compressive and shear stresses are dominant and in the inner substance region where the tissue is avascular. B Biomechanical Studies The linear biphasic theory was applied to evaluate the biomechanical behavior of the type I collagen–based template supported regenerated tissue, fibrin clot supported regenerated tissue, unfilled control, and the normal meniscus tissue. As described before, two compressive biomechanical parameters were obtained by applying the biphasic theory: the aggregate modulus (Ha) and the apparent permeability (k). Prior to analyzing the repaired canine menisci in the die punch model, five normal canine medial meniscus specimens were obtained from skeletally mature mongrel dogs. Two regions (anterior and posterior) of each meniscus were biomechanically tested to establish baseline “normal” values. Normal and experimental menisci were tested using the same procedure described as follows. Twelve experimental canine medial menisci each containing two defect sites were tested. One site was in the anterior region and one site was in the posterior region. The explanted menisci were stored in a freezer (20°C) until biomechanically tested. A 4-mmdiameter specimen was cored out from the repaired site using a disposable biopsy punch. The cylindrical specimen was harvested perpendicular to the tibial surface of the meniscus. The specimen was prepared and tested as described by Mow [33]. C Biochemical Studies Sulfate glycosaminoglycan determination: The sulfate glycosaminoglycans were determined by the method of dimethylmethylene blue (DMMB) assay [63]. Chondroitin 6-sulfate was used for the preparation of the standards. Standards 2 to 12 g/mL were prepared in 0.5 M sodium acetate, pH 6.8. Water content determination: Aliquots of wet specimens were first weighed, and then freeze-dried. The dry weight of specimens was obtained by weighing. The water content was expressed as the weight of water per weight of dry tissue. D Histological Studies Specimens were prepared by standard methods. Specimens were stained with hematoxylin and eosin (H&E), and saffranin-O for visualization of cells, collagen, and proteoglycans. VI RESULTS A Design and Development The purified bovine tendon collagen fibers contained a minimal amount of noncollagenous moieties, as determined by the amino acid composition, hexosamine content, hydroxyproline content, and neutral sugar content of the purified collagen material. Table 1 summarizes the results of the amino acid analysis of three separate preparations of the purified tendon collagen. The amino acid composition of the purified tendon collagen was typical of that reported for type I collagen from bovine and human [64, 65]. The results of chemical

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Table 1 Results of Amino Acid Analysis of Purified Collagen Fibers Residues per 1000 residues Amino acid Aspartic acid Glutamic acid Hydroxyproline Serine Glycine Histidine Arginine Thereonine Alanine Proline Tyrosine Valine Methionine Cystine Isoleucine Leucine Hydroxylysine Phenylalanine Lysine a

Preparation 1

Preparation 2

Preparation 3

Literaturea

43.0 70.5 99.8 31.5 333.3 3.6 50.0 14.4 111.6 121.6 2.7 20.7 4.1 0.0 16.3 25.6 8.7 13.8 28.3

42.6 71.8 100.5 30.7 333.0 3.7 47.8 45.2 108.8 126.1 3.1 20.7 5.8 0.0 18.1 23.4 10.3 13.3 24.5

43.4 75.4 94.5 31.1 330.3 3.7 50.1 15.0 115.5 124.6 3.0 20.3 6.3 0.0 15.3 25.3 10.6 13.9 26.2

49.3 75.4 96.3 30.7 334.4 4.5 48.4 19.0 102.6 122.0 4.6 22.9 6.0 0.0 12.4 25.2 9.4 14.2 29.3

From Ref. 64. Reported values are average values of ten different determinations from tendon tissues.

analysis are summarized in Table 2. The amount of hexosamine in the preparation was less than 0.03%, indicating a minimal contamination of glycoproteins and glycosaminoglycans in the preparation. The hydroxyproline content of the preparation was 13%, which was consistent with the amino acid composition data reported for purified type I collagen from tendon origin. The amount of neutral sugars in the type I collagen is about 0.5% [66]. However, the neutral sugars in the purified collagen have been reduced to a level of 0.01%, indicating the absence of neutral sugar moieties. The neutral sugars that are O-glycosidically linked to hydroxylysine residues were likely to be lost during the base extraction in the purification steps. The hyaluronic acid and chondroitin sulfate contained a nondetectable amount of protein moieties as determined from the results of the microprotein analysis. This finding is consistent with the technical data supplied from the GAG suppliers. Table 2 Results of Chemical Analysis of Purified Collagen Fibers Chemical analysis Hexosaminea Hydroxyproline Neutral sugar a

Number of samples

Results  S.D

25 9 3

0.027  0.005 13.0  0.4 0.008  0.002

Hexosamine content for Avitene is 0.093  0.004 and for Helistat is 0.04  0.002.

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d

Figure 1 The effect of water on the structural parameters of the type I collagen–based meniscus template at different dose level of -irradiation. (a) the effect on the hydrothermal shrinkage temperature; (b) the effect on the mechanical suture pull out strength; (c) the effect on the swelling properties; (d) the effect on the trypsin susceptability.

Figure 1 summarizes the effect of water on the structural parameters at different dose levels of gamma irradiation. In general, gamma had a structural destabilizing effect on collagen molecules and fibers. It was observed that the destabilizing effect was a function of the dose level of gamma irradiation. That is, the higher the dose, the greater the destabilizing effect exerted on collagen. The destabilizing effect was particularly pronounced when the dry matrix was gamma irradiated. The effect on hydrothermal shrinkage temperature, suture pullout strength, and enzyme susceptibility could be clearly seen from the figures. A drop in hydrothermal stability of 11°C was seen between the control and the gamma-irradiated matrix at a dose of 18.7–21.1 kGy (Fig. 1a), a dose level was used for the sterilization of the template. Suture pullout strength was also significantly reduced by the gamma

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Figure 2 Physical appearance of the engineered type I collagen–based meniscus template. irradiation (Fig. 1b). A substantial quantity of the collagen molecules was damaged, presumably by backbone chain scission, since these damaged molecules were not stable at 37°C and were denatured. The denatured portion of the matrix was therefore susceptible to trypsin digestion, as shown in Fig. 1d. The gamma irradiation had a minimal effect on swelling properties since the chain scission likely did not change the net charge of the matrix (Fig. 1c). It is of interest to note that at the dose level of 3.2 kGy, there was no detectable destabilizing effect of gamma on the structural parameters. The incorporation of water molecules to the collagen matrix significantly increased the stability of the matrix to gamma irradiation. There was only about a 4–5°C decrease in hydrothermal stability and about 15% reduction in suture pullout strength when the water content in the matrix reached a level of 30–40% (weight of water per dry weight of collagen matrix) at a dose of 18.7–21.1 kGy. The structure stabilizing effect could also be seen in the trypsin digestion assay (only 2-day digestion data were presented), which showed a drastic reduction in trypsin digestible collagen when the collagen matrix was irradiated in the presence of water at a level 40% or higher. Figure 2 shows the appearance of the engineered type I collagen–based meniscus regeneration template. The template has a physical shape and size similar to the subtotal size (80% of the width) of human meniscus. The template has a length of about 4 cm (anterior to posterior horn), a width of about 0.7 cm (0.2–0.3 cm from the outer edge of the meniscus to inner rim), and a height of 0.45 cm at the peripheral rim; the height drops to 0.1 cm at the inner rim forming a slight concave curvature. Table 3 summarizes the results of characterization studies of the finished, gammasterilized type I collagen–based template. The average apparent density of the dry collaTable 3 Characteristics of the Type I Collagen–Based Meniscus Template Characterization 3

Apparent density (g/cm ) Pore structure (m) Suture pullout strength (kg) Extent of hydration (g/g) Hydrothermal shrinkage temperature (C) Permeability to BSA (% interstitial space) Compressive aggregate modulus (MPa) Compressive permeability (m4/Ns)

Number of samples

Results  S.D.

9 3 41 38 39 10 6 6

0.20  0.02 95%, within 50–400 m 2.36  0.50 3.6  0.5 64.4  2.1 98  5 0.18  0.01 (1120  156)  1015

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Figure 3 Scanning electron micrograph of a type I collagen–based meniscus template in cross-section. 20.

gen–based meniscus template is 0.20 g/cm3. This means that 0.86 cm3 interstitial space per 1 cm3 matrix is not occupied and represents a quite open structure for cell infiltration. The open structure is manifested by the scanning electron micrographs shown in Fig. 3. The perpendicular cross-section (parallel to the load in vivo) shows many dense fibers that are intermingled with an open pore structure of various dimensions, generally from about 50 to about 400 m of irregular shapes with a majority of pores in the range of from 75 to about 300 m. The pore size distribution of the collagen meniscus templates are shown in Fig. 4, where 95% of the pores measured are within the range of from 50 to 400 m. The average suture pullout strength of the hydrated collagen–based template matrix exceeded 2.4 kg. This strength is significantly higher than the minimal strength required for suturing the collagen-based template with the host tissue. Suture pullout strength of about

Figure 4 The pore size distribution of the type I collagen–based meniscus template, calculated based on three preparations.

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1 kg has been proven adequate to insert the template and suture the implant in place in simulated arthroscopically assisted surgeries in human cadaveric knees. The extent of hydration of the matrix was 3.6  0.5 cm3/g matrix. This compares well with the average empty volume (intrafibrillar and interfibrillar) per gram of the matrix, 3.8 cm3/g matrix, calculated from the density data, excluding the volume occupied by the collagen molecules and taking the specific volume for collagen to be 0.7 cm3/g [25]. In other words, the hydrated matrix did not swell to any significant degree. The average time required to fully hydrate a collagen-based template matrix was in the order of 1 min, indicating a high degree of hydrophilicity of the material. The porous and hydrophilic nature of the matrix makes the available interstitial space readily accessible to ions, macromolecules, and other necessary nutrients for cellular metabolism. Table 4 summarizes the observed volume accessible to BSA. The observed volume is consistent with total accessibility of the interstitial space of a porous matrix to macromolecules the size of BSA (MW 67,000). The crosslinked collagen-based template matrix had an average hydrothermal shrinkage temperature of 64°C. The results from the initial pilot study in canines showed that the majority of the collagen from the implant having a hydrothermal shrinkage temperature in the range of 60 to 70°C was resorbed within the first 3 months of implantation, accompanied by ingrowth and new fibrocartilage-like tissue deposition [39,46]. The results of canine study presented subsequently indicated that a substantial amount of new tissue regeneration had already occurred after 12 weeks of implantation with significant matrix degradation. By 6 months, the majority of the template implant had resorbed and replaced by the regenerated tissue. Taken together, it appears that resorption and regeneration are proceeding hand in hand. The compressive stiffness parameter, aggregate modulus (Ha), of 0.18 MPa was similar to the normal canine meniscus (0.15 MPa), indicating that the compressive property is suitable for the physiological function of the template. The permeability (k) for the collagen template, 1120  1015 m4/Ns, was about 300 times the normal canine meniscus (3.71  1015 m4/Ns), indicating that the collagen template is a highly porous matrix for fluid diffusion in and out of the matrix under stress–relaxation cycles. Table 4 Accessibility of Interstitial Space of the Type I Collagen–Based Template to Bovine Serum Albumen Sample no.

Weight of dry sample (g)

Vt (mL/g)

Vobs. a (mL/g)

Vtheo. a (mL/g)a

theo. Vobs. a /Va  100

1 2 3 4 5 6 7 8 9 10

0.0753 0.0742 0.0599 0.0527 0.0741 0.0702 0.1420 0.1158 0.1354 0.1540

4.49 4.47 5.29 5.04 4.36 4.06 4.97 5.13 4.83 5.22

3.21 3.13 3.84 3.91 3.13 3.13 3.61 3.89 3.69 4.22

3.35 3.33 4.15 3.91 3.22 2.92 3.83 3.99 3.69 4.08

96 94 93 100 97 107 94 98 100 103

a

Vtheo.  Vt  Vi , where Vi represents that part of the volume which is inaccessible to BSA. Vi of 1.14 mL/g was a used based on Ref. 42.

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B Biocompatibility Studies 1 Humoral Immune Response All positive control rabbits mounted an immune response to the type I collagen–based template implant, with antibody levels rising over time until a maximal response was reached. Figure 5 shows a typical humoral immune response of antibody production in a positive control rabbit at a series of serum dilutions. The maximum antibody production was reached at approximately 7 weeks. There was no additional increase in the production of antibodies from 7 weeks to 16 weeks despite the additional boosters given at 10 and 13 weeks postimplantation. Figure 6 summarizes the ELISA test of antibody production in positive control and implant rabbits at a 1:10 serum dilution as a function of implantation time. No humoral immune response to the collagen implant was demonstrated. 2 Cell-Mediated Immune Response a. Spleen Cell Response to Type I Collagen–Based Template. Figure 7 shows the response of spleen cells to collagen-based template implant. The only significant differences (p 0.05) were in the positive control group (pulverized collagen template injected with Freund’s adjuvant) at 2 weeks and 6 months posttreatment. This expected response indicated that the mice had the ability to respond to the collagen-based template. Thus, the type I collagen–based template did not cause a statistically significant sensitization of the mouse spleen cells to type I collagen–based template materials.

Figure 5 Positive control rabbit injected with pulverized type I collagen–based meniscus template in Freund’s adjuvant over a period of 4 months. This rabbit shows a response typical of that seen in the four positive control rabbits.

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Figure 6 ELISA of rabbits for antibodies to type I collagen–based meniscus template at serum dilution 1:10.

b. Lymph Node Cells Response to Type I Collagen–Based Template. Figure 8 shows the response of lymph node cells to type I collagen–based template implant. The only significant differences were seen in the 2-week groups of mice. All three (implant, pulverized implant, pulverized implant with Freund’s adjuvant) were significantly increased only because of the low sham values for that group. At 2 and 6 months, there were no significant differences among the test groups.

Figure 7 Plot of proliferation stimulation index of mice spleen cells to type I collagen–based meniscus template against various treatment conditions of antigens at different times of implantation.

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Figure 8 Plot of proliferation stimulation index of mice lymph cells to type I collagen–based meniscus template against various treatment conditions of antigens at different times of implantation.

3 Additional Biocompatibility Studies The results of the additional biocompatibility tests show that the type I collagen–based template is noncytotoxic, nonpyrogenic, and nonhemolytic. No mutagenic response was observed for collagen, GAGs, or from the trace amounts of chemicals left in the device during material processing. Thus, the collagen-based template developed can be used for in vivo implantation. C Animal Studies 1 Die Punch Model Studies Figure 9 shows typical explants of the medial menisci after 3 and 6 months of collagen template implantation. The location of the template could still be seen at the 3-months time point. However, at 6 months, the demarcation between the template and the host becomes

a

b

Figure 9 Gross appearance of explants of the medial menisci of the die punch model. (a) after 3 months implantation of type I collagen–based meniscus template; (b) after 6 months implantation of type I collagen–based meniscus template.

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Figure 10 Histological view of a 3-month type I collagen–based meniscus template die punch explant (100).

less defined, indicating that the regenerated tissue was fully integrated with the host meniscus. The fibrin clot–filled lesions looked generally similar to the collagen-based template–supported regeneration. The unfilled control lesions were filled with tissue that was more fibrous in appearance. Figure 10 shows a histologic section of a collagen template explant at 3 months. There was an extensive cellular infiltration and new tissue regeneration replacing most of the template material which had been resorbed. The cells were viable and had physical characteristics resembling those of fibrochondrocytes. At the template–host tissue junction, there was firm integration between the template and the host meniscal tissue. The fibrin clot–regenerated tissue and unfilled controls were also integrated with surrounding host tissue. However, the new collagen fibers appeared smaller and more chondroid than did the collagen template–regenerated tissue. The regenerated tissues within unfilled defects were likely due to the blood clot from the hemarthrosis that resulted postoperatively. At 6 months (Fig. 11a), residuals of collagen template could still be seen, as evidenced by the dark staining material. The cells appeared viable and abundant. The general physical appearance of the cells were more mature than the cells at 3 months and were similar to those of normal meniscal fibrochondrocytes. The newly regenerated collagen appeared to be lining up and was more organized than that seen at 3-month time point. At this more advanced regeneration stage, ample glycosaminoglycans staining (Fig. 11b) could be seen in the collagen template implant indicating the synthesis of proteoglycans by the infiltrated cells. The fibrin-regenerated tissue was also taking on a more organized appearance, but it was less dense with smaller collagen fibers than the collagen template tissue.

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b

Figure 11 (a) Histological view of a 6-month type I collagen–based meniscus template die punch explant (25). (b) 6-month type I collagen–based meniscus template die punch explant, stained with Saffranin-O for GAG (100).

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Tissue from the unfilled controls was beginning to take on a more organized appearance, but it was still chondroid with small collagen fibers. 2 Biomechanical Studies Figure 12 illustrates the biomechanical properties of the three treatments at two time points and the normal meniscus tissue. Normal control canine specimen values were averages of the anterior and posterior regions, as no statistical difference was found between regions. A one-way ANOVA (p0.05) was conducted for both time groups and both biomechanical parameters. Though a Student t test indicated no statistical differences in aggregate modulus or permeability between repair treatments and normal, the regenerated tissue from the collagen template exhibited biomechanical characteristics most similar to normal tissue for both 3-month and 6-month groups. An overall increase in compressive stiffness and decrease in permeability was found at 6 months compared to 3 months.

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Figure 12 Results of biomechanical studies of the three treatments and normal meniscus tissue of the die punch model at 3 and 6 months: (a) aggregate modulus (Ha); (b) apparent permeability (k).

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Table 5 Results of Biochemical Studies of the Explants 3 months Type of specimen Normal meniscus Collagen template regenerated tissue

6 months

No. of samples

Water content (%)

GAG content (%)

No. of samples

Water content (%)

GAG content (%)

6

66.6  1.35

2.96  1.19

6

66.6  1.35

2.96  1.19

6

77.9  2.26

1.25  0.50

6

77.1  5.32

2.81  1.60

Note: Results expressed as mean  standard deviation.

3 Biochemical Studies The results of the water and glycosaminoglycan content determinations are shown in Table 5. The water content of the normal meniscus was 67%. The water content in the normal meniscus tissue was consistent as seen from a small standard deviation of the data. This water content was comparable to that previously reported for the canine meniscus [67]. The water content in the regenerated tissue at the 3-month time point was 11% higher than the normal meniscus, or 78%. The water uptake capacity for the nonimplanted collagen template was 3.6 g water per gram of dry matrix (Table 3). That is, the water content in the hydrated nonimplanted matrix is 78%. It appeared that the original template material had been replaced by the newly synthesized matrix. This level of water content did not change at 6 months, suggesting that the structure of the regenerated matrix is still at a less organized, nonmature state. The GAG content of the normal meniscus was 2.96% of the wet tissue. This value is comparable to that reported previously [68]. The average GAG content of the regenerated tissue at 3 months was 1.25%, and it increased to 2.81% at 6 months, a level approaching the GAG content in the normal meniscus. D Subtotal Resection Model Figure 13 shows the gross appearance of explants at 3 weeks and 6 months postoperation. In general, there was excellent healing of the matrix template and the associated regenerated tissue with the host meniscus tissue. The template implants were firmly attached to the

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Figure 13 Gross appearance of type I collagen–based meniscus template explants of subtotal resection model at (a) 3 weeks and (b) 6 months.

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b

Figure 14 Histological view of type I collagen–based meniscus template explants of subtotal resection model at (a) 3 months and (b) 6 months (100). host at all time points. In addition, there was an increase in tissue integration with time at the outer periphery that appeared to progress toward the inner rim of the template implant. Figure 14 shows histologic specimens at 3 and 6 months postoperation. At 3 months postoperation, much of the template implant had been resorbed and was replaced by the regenerated tissue (Fig. 14a). The collagen fibers were lining up along the lines of stress. Many areas had the appearance of fibrocartilage with round-shaped cells consistent with chondrocytes. The host–implant junction was dense and well integrated. The regenerated tissue overall was still immature. A significant amount of template implant still remained and could be easily identified by the H&E staining. At 6 months (Fig. 14b), the specimen was consistent with ongoing maturation of the new tissue. The integration of the host with the implant was excellent. However, a small amount of implant collagen still remained in some specimens. A significant amount of dense fibrocartilage-like tissue was formed in various areas throughout the specimen. The collagen bundles were aligned, and in some specimens the cells were in lacunae. VII DISCUSSION Crescent-shaped menisci of the knee joint serve important biomechanical functions. Damage to the meniscus can affect the normal function of the knee joint. Partial removal of the damaged meniscus is the most common surgical treatment for a torn meniscus. It is now known that removal of the meniscus frequently leads to later development of degenerative arthritis. Consequently, a variety of replacement materials have been sought. At the present time, none of the permanent meniscus substitutes has proven effective. It is generally recognized that mismatch of the mechanical properties between the substitute and the host meniscus and the material fatigue and degradation of the substitute are the major drawbacks for permanent meniscus replacement. Meniscus allotransplantation [69–71] has been used for the replacement of irreparably damaged meniscus. Although acceptable to many surgeons, many questions remain unanswered. Many investigators are continuing to examine changes in biomechanical characteristics of the allograft with time, incongruities of the size and shape, insertion and preservation techniques, control of disease transmission, and the graft shrinkage after implantation [71]. Thus, the long-term efficacy of meniscus allografts remains to be determined. As the science and technology advance in the field of tissue engineering, it is apparent that innovative approaches must be developed to replace the traditional methods of tis-

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sue repair. Tissue engineering is an emerging medical science and technology involving three key components: extracellular matrix, biological signal, and cell. Since a tissue engineering approach for orthopedic tissue repair and regeneration is a new branch of medical science, we would like to offer the reader our thoughts of tissue engineering as a solution to meniscus repair and regeneration. Our working hypothesis for meniscus regeneration is that the meniscus has the intrinsic ability to regenerate provided that the biological environment is suitable for regeneration. Thus, in order for any tissue engineering product to function in a load-bearing application, one must first be certain that the design requirements for that product are met. Among the three components of a tissue engineering product for meniscus regeneration, we consider the extracellular matrix to be the most critical element of the three, since the extracellular matrix serves the biomechanical function of the meniscus. Thus the initial biomechanical properties of the extracellular matrix and the subsequent biomechanical properties of the regenerated and remodeled tissue must be adequate for the device to function. In addition, the extracellular matrix must be conductive to cells as well as permeable to nutrients. The biological signal and the cell, the other two tissue engineering components, can accelerate the overall regeneration and remodeling process and may provide a more ideal biological environment for cellular infiltration and matrix synthesis. Thus, the overall goal of the present investigation was to design and engineer a resorbable extracellular matrix template and to test the working hypothesis of the template for the support and guide of meniscus tissue regeneration in a canine model. In this stage of development of a tissue engineering product for meniscus tissue regeneration, we have taken the view that without a well-designed extracellular matrix, a tissue engineering product cannot be functional. The successful development of a functional extracellular matrix template would represent a major achievement toward the development of an ultimate tissue engineering product for meniscus repair and regeneration. The concept of a porous resorbable template for supporting meniscus regeneration was recently tested in vivo. Indeed, the initial porous collagen-based template could serve as a template for meniscus regeneration in dogs and in humans [39,46,72]. However, several areas have been identified which require further investigation in terms of design requirements of a resorbable matrix to function as an ideal template for meniscal regeneration. In this chapter, we have reviewed the background and the rationale of the design requirements of a resorbable type I collagen–based meniscus template. We will further discuss the key requirements here. We consider the biocompatibility of the materials and their degraded products a prerequisite for resorbable, implantable device development. The results of biocompatibility and preliminary clinical studies clearly show that the engineered collagen template is safe for human implantation. The development of an engineered collagen-based template requires that the template implant be sterile at a sterility assurance level (SAL) of 106. Among the acceptable sterilization methods available, we have developed a gamma sterilization method for the collagen-based meniscal template. Since gamma irradiation has a structural destabilizing effect on collagen, the structural stabilization with water molecules during sterilization has provided a way to sterilize the template implant with minimal structural destabilization on collagen. This is particularly important in orthopedic tissue regeneration applications where load bearing of the template implant is required for in vivo function. Without the water stabilization on the collagen structure during gamma irradiation, the resulting implant would not fulfill the initial design requirements for meniscus regeneration applications.

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One of the major requirements of using purified, reconstituted, type I collagen–based template for load-bearing applications is the mechanical properties of the template implant. In the pilot study in canines using the initial prototype, the mechanical failure may have accounted for 25% of the total implants [46]. The excessive mechanical forces may have resulted in compressive and shear stress induced dislodgment of the implant at the suture line, which connects the implant to the peripheral meniscus rim, before the repopulation of the cells and deposition of the new tissue occurred. Compressive and shear stresses may also have resulted in the fraying of the matrix material, causing cleavages and delamination of the implant. These observations have resulted in our new design hypothesis, i.e., an accurately fit device would minimize biomechanical destruction and improve the effectiveness of the template implant for supporting and guiding meniscal tissue regeneration. There are other key mechanical property considerations for resorbable templates for load-bearing tissue regeneration applications, including the initial mechanical strength of the template, the mechanical characteristics of the regenerated tissue/template composite during the regeneration and healing period, and the mechanical characteristics of the regenerated meniscus. The adequate initial mechanical strength permits the insertion of the template to be performed with an arthroscopically assisted approach as well as by conventional open procedures and also provides the device with in vivo stability after implantation. The mechanical strength of the template is mainly controlled by the density of collagen material in the matrix and by the extent and the uniformity of chemical crosslinking. The density of a collagen matrix must be optimized to the extent such that the interfibrillar distances are large enough for cell infiltration. An interfibrillar distance greater than 50 m should be sufficiently large since the cells in general have a size between 10 to 50 m [26]. However, if the interfibrillar distance is less than 10 m, which prevents cell infiltration, the rate of resorption will more likely precede the rate of new tissue deposition. Such an implant may be strong initially, but as the implant resorbs the extent of reduction in mechanical strength of the implant may not be sufficiently compensated by the new tissue laid down on the surface of the implant. In other words, the mechanical failure may precede the regeneration, resulting in implant failure. On the contrary, if the interfibrillar distance is too large, then the material may not be strong enough to sustain suturing and provide adequate initial mechanical support for the device to function. The geometrical dimension of the template provides a guide or boundary for cellular regeneration. The geometrical factors for meniscus regeneration are controlled by the dimension of the template as well as by the in vivo conditions to which the template is subjected. The physical dimension of the template under in vivo conditions provides a guide for cells to repopulate within the boundary defined by the template. Thus, a simulation of the anatomical structure of a template will replicate the natural dimensions of the meniscus. The final anatomical form of the regenerated meniscus is defined by the mechanical forces that act upon the template and the space available between the femoral condyle and tibial plateau. Hydrophilicity and swelling both facilitate the diffusion of nutrients into the interstitial space. This is of particular importance in tissues where nutrient supply is largely dependent upon the mechanism of diffusion rather than through vascular channels. The survival of meniscal fibrochondrocytes of the acellular midmaterial is dependent on the physicochemical properties of the matrix for nutrient diffusion [34]. The accessibility of ions, peptides, growth factors, and macromolecules permits the cells to perform their differentiated functions. The complete accessibility of the interstitial

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space to test molecule of molecular weight of 67,000 daltons indicated the adequacy of the permeability of the matrix. The results of immunological studies of the type I collagen–based template implants in rabbits and mice have shown that the collagen fibers produced by the purification steps and the subsequent engineering and chemical crosslinking did not alter the original hypoimmunogenicity of the collagen material. The type I collagen–based template implant has very little, if any, impact on the immune system. Overall, from an immunological point of view, the type I collagen–based template implant is safe for human implantation. The other battery of tests performed on the type I collagen–based template implant all show that the implant is biocompatible and is safe for human implantation. The characteristics of the template (Table 3) which we targeted in designing the template have been given a rigorous test in dog models. The dog was allowed to ambulate and bear weight immediately after surgery. The results from the die punch model study have provided important information of the continued remodeling of the regenerated tissue toward meniscus-like tissue. At 3 months implantation, a majority of the template had already resorbed. The infiltrated cells have physical appearance similar to fibrochondrocytes. As the template resorbs, the new tissue was laid down in accordance with the dimension of the template. At 6 months, most of the template implant had resorbed. The cells at this time point resembled the morphology of fibrochondrocytes and ample GAG-containing materials were synthesized in the matrix. The collagen fibers had also started to line up to provide the necessary biomechanical properties of the regenerated tissue. The results of biomechanical studies confirm the design requirement that the initial template has compressive stiffness similar to the native meniscus tissue and that the template is highly porous (high permeability) for cellular ingrowth and nutrient exchange. At 6 months after implantation of the template, the biomechanical properties were maintained at the level suitable for in vivo function. The regenerated tissue from the type I collagen–based template showed superior biomechanical characteristics compared with the regenerated tissues from the fibrin clot and from the defect that was left unfilled. However, there were several limitations in the biomechanical study that must be considered when evaluating the data. The sample size of each defect group was too small to eliminate variations among the individual dogs. Poor quality of some of the specimens prepared caused difficulty when harvesting a 4-mm plug. A 4-mm repaired defect may not be large enough to induce a regenerative response from which consistent biomechanical data can be obtained. For example, when harvesting a specimen from the repaired region of a meniscus, it was often difficult to obtain a 4-mm disc that consisted of pure regenerated tissue. The creation of two defects within one meniscus may not produce an isolated biomechanical response from each defect. A larger, individual meniscus defect may be required to generate more consistent and indicative of the behavior of the regenerated tissue in vivo. In addition, long-term biomechanical capabilities of the regenerated tissue are required to assess its long-term durability. Further studies are required to obtain more definitive results of the biomechanical characteristics of the regenerated tissue. The regenerated tissue from the type I collagen–based template has a water content slightly higher than the normal meniscus even at the 6-month time point, suggesting that the regenerated tissue is still in the maturation process. This observation is consistent with the histologic observations that the collagen fibers are lining up toward a more organized structure as seen in the normal meniscus. Importantly, the infiltrated cells synthesize glycosaminoglycans, presumably in the form of proteoglycans that are commonly associated with cartilagenous tissues. This finding is also consistent with the histologic observation

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based on Saffranin-O staining. The content of the GAGs in the regenerated tissue approaches that of the normal meniscus at 6 months after implantation, suggesting that cells are functioning as fibrochondrocytes. The results from gross appearance and histologic evaluations in the subtotal resection model have demonstrated that the prototypes engineered according to the design requirements can satisfy not only the requirements for in vivo function as a template to support the meniscus regeneration while functioning as a temporary meniscus substitute, but also for supporting continued remodeling of the regenerated tissue toward fibrocartilage-like tissue. The results from the dog studies demonstrate that the design requirements of the engineered extracellular matrix template have been met and that the template can serve as a cell conductive matrix to support and guide the meniscus tissue regeneration. The successful outcome from these studies have led to the FDA approved clinical studies of the type I collagen–based extracellular matrix template for meniscus preservation and restoration. The results of preliminary clinical studies of this type I collagen–based matrix template have been published elsewhere [73,74]. The successful design and engineering of an extracellular matrix has laid the foundation for the tissue engineering approach to meniscus tissue regeneration. As the field of biological signals and cell biology advances, we will be able to apply the scientific information from these areas to design and engineer a tissuelike matrix that has a biological environment that can accelerate the regeneration as well as the maturation process, i.e., a bioactive matrix that is both cell conductive and inductive for tissue regeneration. For example, recent advances in growth factor research have resulted in several clinical studies of the growth factors as biological signals for tissue wound healing applications. Our thoughts in this area are that if a functional extracellular matrix contains a biological signal that can recruit the local undifferentiated mesenchymal cells to the repair site and induce them to differentiate into functional cells of the tissue of interest, the regeneration process will be enhanced. Several candidate growth factors may be useful for potential meniscus tissue regeneration applications such as the bone morphogenetic proteins (BMPs) and basic fibroblast growth factor (bFGF). Bone morphogenetic protein-2 has been shown in vitro to be chondrogenic [75] and bFGF is a potent angiogenic factor [76]. Both growth factors can potentially accelerate the healing process by enhancing the regeneration of the desirable functional tissues. Advances in the field of cell biology have also pushed the research of tissue engineering products to a new level. The most noted is in articular cartilage repair employing either the differentiated articular chondrocytes [77–79] or the undifferentiated mesenchymal stem cells [80,81]. The cells are delivered to the injured or diseased articular cartilage site for the cells to differentiate (for stem cells) or maintain the differentiated state (for articular chondrocytes), and synthesize articular cartilage matrix. Our view in this area is consistent with our general view of tissue regeneration by a tissue engineering approach. We believe that in order for cells to function, one must provide a functional extracellular matrix that meets the design requirements for articular cartilage regeneration. This concept is also directly applicable to the meniscus regeneration. That is, a functional matrix which contains the cellular component will enhance meniscus regeneration by virtue of pre-existence of the cells within the matrix to perform their function. It should be emphasized that the design parameters discussed here are not independent. Thus, in designing a medical implant, particularly a resorbable template implant for regeneration applications, the design parameters must be considered as a whole. Our goal

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in designing a meniscus regeneration template is analogous to fitting a set of interdependent parameters of an equation to an experimentally observed curve by a least squares minimization approach (through iteration) such that a convergence to the true minimum of the sum of the squares of residuals is obtained [82]. In doing so, a set of initial input based on the best information available on each of the design parameters is first tried. The ultimate converging factor here is the best meniscus template to achieve a full regeneration of the functional meniscus. Based on the information collected thus far, our best input for a meniscus regeneration template at present is that shown in Table 3. The results of biocompatibility and the in vivo dog studies which we presented in this paper demonstrate that the designed prototype fulfils the requirements as a template for supporting and guiding the meniscus tissue regeneration in the knee joint. As we move forward in the field of tissue engineering, an ultimate meniscus-like tissue (the final input) can be engineered in vitro. Such a meniscus-like tissue should have the best chance to replace the missing meniscus in those patients who suffer from severe joint degeneration. ACKNOWLEDGMENTS We thank Dr. Steven P. Arnoczky for conducting animal implantation and performing histologic analysis of the explants, Drs. Jun-Kyo Suh and Savio L.-Y. Woo for performing the compressive mechanical testing of the collagen meniscus implants and explants, Dr. Cahir A. McDevitt for performing the water content and glycosaminoglycan analyses of the explants, and Dr. Judith B. Ulreich for conducting the immunological studies. REFERENCES 1. Li S. T. 2000. Biologic biomaterials: tissue-derived biomaterials (collagen). In: The Biomedical Engineering Handbook, Bronzino J. D. Ed. CRC Press, Boca Raton, FL, pp. 42.1–42.23. 2. Fithian D. C., Kelly M. A., Mow V. C. 1990. Material properties and structure–function relationships in the menisci. Clin. Orthop. 252:19–31. 3. Arnoczky S. P., Adams M., Dehaven K., Eyre D., Mow V., Kelly M. A., Proctor C. S. 1991. Meniscus. In: Injury and Repair of the Musculoskeletal Soft Tissues. Woo, S. L.-Y., Buckwalter J. A., Eds. Am. Acad. Orthop. Surgeons: Park Ridge, IL, pp. 483–537. 4. Dehaven K. E., Arnoczky S. P. 1994. Meniscal repair. J. Bone Joint Surg. 76A:140–152. 5. Newman A. P., Daniels A. U., Barks R. T., 1993. Principles and decision making in meniscal surgery. J. Arthrosc. Rel. Surg. 9:33–51. 6. Fairbank T. J. 1948. Knee joint changes after meniscectomy. J. Bone Joint Surg. 30B:664–670. 7. Jackson J. P. 1968. Degenerative changes in the knee after meniscectomy. Br. Med. J. 2:525– 527. 8. Lutfi A. M. 1975. Morphological changes in the articular cartilage after meniscectomy: an experimental study in the monkey. J. Bone Joint Surg. 57B:525–528. 9. Cox J. S., Nye C. E., Schaefer W. W., et al. 1975. The degenerative effects of partial and total resection of the medial meniscus in dogs’ knees. Clin. Orthop. 109:178–183. 10. Lanzer W. L., Koneda G. Y. 1990. Changes in articular cartilage after meniscectomy. Clin. Orthop. 252:41–48. 11. Hede A., Larsen E., Sandberg H. 1992. Partial versus total meniscectomy. J. Bone Joint Surg. 74B:118–121. 12. Timpl R. 1982. Antibodies to collagen and procollagen. Methods Enzymol. 82:472–498. 13. Timpl R. 1976. Immunological studies on collagen. In: Biochemistry of Collagen, Ramachandran G. N., Reddi A. H., Eds. Plenum: New York, pp. 319–375.

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14 Biomaterials for Cartilage Tissue Engineering Hani A. Awad, Geoffrey R. Erickson, and Farshid Guilak Duke University Medical Center, Durham, North Carolina

I INTRODUCTION Articular cartilage is the thin layer of deformable, load-bearing material that lines the bony ends of all diarthroidal joints (Fig. 1). The primary functions of this tissue are to support and distribute forces generated during joint loading and to provide a lubricating surface to prevent wear of the joint. Under normal physiological conditions, articular cartilage can perform these essential biomechanical functions with little damage or wear over the human lifespan. However, articular cartilage is often injured and has a limited capacity for repair. Thus, even minor lesions or injuries may lead to progressive damage and joint degeneration. Several methods have been used clinically to promote cartilage repair in cases of injury and arthritis, albeit with inconsistent or undocumented success [1,2]. One of the potential explanations for the poor repair response of articular cartilage is the lack of a blood supply or a source of undifferentiated cells that can promote repair. For this reason, many clinicians have utilized various means of penetrating the subchondral bone to induce bleeding at the site of the damaged cartilage. These techniques have included drilling, abrasion, and microfracture of the subchondral bone, oftentimes in a manner that penetrates to the marrow cavity [3,4]. Other surgical repair techniques include periosteal and perichondral tissue grafting [5], osteochondral auto- and allografting [6–8], and chondrocyte transplantation (Carticel®, Genzyme Biosurgery, Cambridge, MA) [9,10]. These methods generally lead to the formation of a fibrocartilaginous repair tissue that in some cases is satisfactory and decreases pain and disability [3]. In general, however, fibrocartilaginous repair tissue does not have the same mechanical properties of normal articular cartilage and does not seem to function effectively as a replacement for normal cartilage [11]. 267

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Figure 1 Articular cartilage lining of the knee joint. (From Ref. 151.) Recent advances in biology and engineering have introduced the concept of “tissue engineering,” whereby cells, scaffolds, DNA, proteins, and/or protein fragments are surgically implanted to promote cartilage repair or replacement [12–16]. Tissue engineering merges aspects of engineering and biology. Many rapid achievements in this field have arisen in part from significant advances in cell and molecular biology (e.g., the isolation and manipulation of cells, genes, and growth factors), biomaterials (new and innovative delivery vehicles), and the integration of biology and materials to deliver viable cells in compatible support structures. Despite many advances, at this writing there is only one cell-based repair procedure available for clinical use (Carticel), which involves the isolation and proliferation of autologous chondrocytes ex vivo, followed by reimplantation of cells into the cartilage defect. The defect is then covered by a flap of autologous periosteal tissue [4,9,11]. The long-term success of this technique has not been shown in animal models, although clinical outcomes are initially satisfactory. Many other such techniques are currently being explored and involve a range of synthetic and biological matrices that may be seeded with cells to enhance repair. Such technologies may also be augmented by various growth factors such as transforming growth factor-beta (TGF-), bone morphogenetic protein-6 (BMP-6), basic fibroblast growth factor (bFGF) and insulin-like growth factor (IGF), which have been shown to promote chondrogenic potential [17–19]. It is hoped that the combination of living cells implanted in appropriate biomaterials coupled with the identification of chondrogenic promoting factors will introduce many new cartilage repair options in the near future. In this chapter, the biomaterials that have been used to date for the repair of articular cartilage in cellular and acellular approaches are reviewed. First an overview of the structure and function of articular cartilage is presented in order to lay out the foundation for understanding the biology and biomechanics of normal cartilage. Next, the general criteria for selecting proper biomaterials for successful tissue engineered repair are discussed. Finally, some of the most commonly utilized materials in cartilage tissue engineering are presented.

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II STRUCTURE AND FUNCTION OF ARTICULAR CARTILAGE Recent advances in cartilage research have greatly improved our understanding of the physiology, ultrastructure, and biomechanical properties of articular cartilage in both normal and pathologic conditions. These advances have provided an improved understanding of the etiopathogenesis of joint diseases and offer hope in the development of new clinical modalities for the prevention or treatment of various conditions. Whether they are based on pharmaceutical, surgical, biophysical, or tissue engineering approaches, development of new treatment modalities requires a thorough understanding of the structure, function, and biology of normal and diseased cartilage. A Structure and Biology of Articular Cartilage Articular cartilage consists of an organic extracellular matrix that is saturated with water and a small volume fraction of cells. The water phase of cartilage constitutes from 65 to 85% of the total tissue weight and plays an important role in dictating many physical properties as well as the transport of nutrients, metabolites, and soluble mediators [20–23]. The dominant structural components of the solid matrix are the collagen molecules (~75% by dry tissue weight) and the negatively charged proteoglycans (~20–25% by dry tissue weight). Several other collagen species, smaller proteoglycans, and other proteins are also present in the tissue but at much lower concentrations. The solid matrix of normal articular cartilage has a highly specialized ultrastructure that varies with depth from the cartilage surface to the subchondral bone [24]. In this regard, the tissue may be categorized into a number of successive “zones” based on extracellular matrix structure and composition, as well as cell shape and cell arrangement (Fig. 2) [25,26]. Collagen fibers in the surface zone of cartilage are of small diameter, densely packed, and oriented parallel to the articular surface [27]. This zone is also characterized by a relatively low proteoglycan content and a low permeability to fluid flow [28,29]. In the middle, or transitional, zone, the collagen fibers form an arcade-like structure upon

Figure 2 Articular cartilage is structurally divided into four zones. In the superficial zone, the chondrocytes are flattened and appear to parallel collagen fiber orientation. In the middle zone, chondrocytes are spherical and the collagen fibers form an arcade-like structure upon which a randomly arranged one is superimposed. In the deep zone, the collagen fibers are oriented perpendicular to the bone and the chondrocytes are slightly elongated in a columnar arrangement. The deep zone of the articular cartilage is separated from the calcified cartilage zone by histologically distinguishable “tidemark.” (From Ref. 152.)

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which a randomly arranged one is superimposed [24]. The proteoglycan concentration is highest in this region [30]. In the deep zone, adjacent to the zone of calcified cartilage and subchondral bone, the collagen fibers are larger and form bundles which are oriented perpendicular to the bone, and the proteoglycan content is low [28,30]. Chondrocyte density is highest in the surface zone and decreases with depth [31]. Chondrocyte shape seems to parallel the local collagen fiber orientation with a flattened discoidal shape in the surface zone, is nearly spherical in the middle zone, and is slightly elongated in a columnar arrangement in the deep zone [26]. The zone of calcified cartilage separates the hyaline articular cartilage from the subchondral bone, and the deep zone of the articular cartilage is distinguished from the calcified zone by a thin line termed the tidemark [32–34]. This transition layer of calcified cartilage, of stiffness intermediate to that of articular cartilage and that of the subchondral plate, is hypothesized to serve as a tethering mechanism for collagen fibers[33] and is believed to minimize the stiffness gradient and therefore the shear stresses at the subchondral interface [35]. B Biomechanics of Articular Cartilage In understanding the requirements for utilizing different biomaterials for cartilage repair, it is necessary to examine the normal functional properties of articular cartilage. This tissue functions as a smooth wear-resistant surface facilitating load support and transfer, and provides translation and rotational motions between articulating bones of the skeleton. When an external load is applied to a joint, articular cartilage deforms [36], presumably to increase joint contact areas and ultimately decrease contact stresses. Loading and deformation of cartilage generates a combination of tensile, compressive, and shear stresses within the tissue as well as osmotic and pressure gradients and streaming potentials. The mechanical response of cartilage to these stresses is highly specialized due to the tissue’s unique composition and structural organization. Since the tissue exhibits viscoelastic properties, loading is associated with time-dependent effects such as creep, stress relaxation, and energy dissipation [37–39]. These viscoelastic behaviors arise from both interstitial fluid flow through the porous-permeable solid matrix and from physical interactions between the solid matrix constituents (e.g., collagen and proteoglycan). Cartilage also exhibits highly nonlinear mechanical properties such as strain-dependent moduli [40], strain-dependent hydraulic permeability [41], and a difference of nearly two orders of magnitude in tensile and compressive moduli [42]. These properties are also anisotropic, particularly in tension, and vary significantly depending upon site location within the joint and depth from the tissue surface [43]. Additional mechanical behaviors also include the presence of internal swelling pressures that give rise to inhomogeneous residual adaptation to physical demand [44]. Accordingly, the biomechanics of cartilage has been widely studied using models that account for the multiple phases of cartilage (collagen solid matrix, interstitial fluid, and mobile and “fixed” ionic groups) and their interactions. Such model development, in combination with the experimental testing of the material properties of cartilage, is important in understanding the mechanisms by which cartilage functions within the joint and how degeneration and aging might affect cartilage function. Despite the vast literature on the biomechanical properties of normal and pathological cartilage, it is still to be determined which of these properties need to be replicated for a successful engineered repair. An understanding of the functional properties of normal and repair tissues will likely provide important information guiding the selection of a scaffold biomaterial for cartilage tissue engineering.

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III TISSUE ENGINEERING BIOMATERIALS REQUIREMENTS AND CRITERIA The field of tissue engineering has emerged over the years and is continually evolving to refine the complex interactions between cellular and noncellular components. In this regard, tissue engineering has become a diverse multidisciplinary field where life sciences and engineering meet. When setting out to design a tissue engineered construct, certain criteria based on our understanding of cartilage biology and biomechanics must be sought in order to achieve ultimate functional (biological, biomechanical, and clinical) and commercial success. In this section, we discuss some aspects of the design criteria for the functional success of tissue engineered cartilage. It is generally accepted that the engineered regeneration of articular cartilage will require the controlled interactions of cells, biomaterials, and possibly soluble mediators (i.e., growth factors). However, other environmental factors, such as the biophysical and biological milieu of the construct in vivo, will have an important influence on the metabolic activity and phenotype of the implanted cells. There are several criteria that a biomaterial must meet in order to maximize the chances of a successful repair. For example, one would expect that the material should promote a balanced cellular metabolism and facilitate functional tissue growth. Other criteria include biocompatibility, whether or not the biomaterial should degrade over a specified time period, the biological and biomechanical properties, the ease of handling, and the ability to sterilize and package the material. Furthermore, other criteria may depend on many different factors involving the characteristics of the defect being treated (e.g., size and severity of the lesion) and the nature of the treatment being sought. A Promoting Balanced Cellular Metabolism and Facilitating Functional Tissue Growth In order to obtain a successful tissue engineered construct, the biomaterial scaffold and the living cells must coordinate their existence and interactions in a controlled manner throughout the life of the engineered tissue. In essence, the long-term goal of cartilage tissue engineering is the successful regeneration of the native tissue and the restoration of tissue function to that of the normal, original tissue. Presumably, the regeneration of new tissue must proceed in concert with the degradation of the scaffold material such that the functional properties of the implant do not drastically change from culture to implantation and then subsequent remodeling in vivo. One interpretation of this balance in cellular metabolism controlling synthesis and degradation kinetics is depicted in Fig. 3. Quantitative measures of “functional success” still remain to be defined and universally accepted through a variety of biomechanical and biochemical analyses. Through cellular and tissue explant studies, scientists have found that metabolic activity and phenotypic expression of chondrocytes are strongly influenced by specific biological and biophysical environmental factors [45]. Therefore, it is imperative that these environmental factors are replicated or at least precisely controlled around the cells being used in a tissue engineered cartilage construct to provide the optimal biological and biophysical environmental conditions for cell growth and function. For example, chondrocytes embedded in the cartilage extracellular matrix are exposed to a complex set of biological factors and biophysical stresses during physiological static and dynamic loading conditions. The biological factors include soluble mediators such as growth factors and cytokines, and the biomechanical factors include deformation, fluid flow, and varying osmotic pressures and oxygen tension. However, despite our understanding of these factors, there is limited quan-

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Figure 3 Balance in synthesis and degradation kinetics in tissue engineered cartilage constructs is essential in maintaining tissue engineered cartilage functionality.

titative information available on the various stimuli (and their interactions) to which chondrocytes are exposed. In this respect, studies must utilize novel biochemical techniques as well as physical and theoretical models in order to attempt to approximate the chemical and biomechanical environment around chondrocytes within a normal joint during loading [46]. Recent studies have focused on elucidating the role of physical attachment on the extracellular matrix [47,48] and suggest such cell–matrix interactions alone may have separate critical signaling cascades associated with them. Maintaining such cell–matrix interactions may be crucial to the success of cartilage tissue engineering. In summary, many investigators now believe that in order for a scaffold biomaterial to promote balanced cellular metabolism and facilitate functional tissue growth, it must be constructed in a manner that mimics the natural environment to which the chondrocytes would be exposed in healthy tissue. Because to date no single biomaterial exists that seems to replicate the exact in vivo environment of chondrocytes, it will be necessary to prioritize the external biological and biomechanical factors and include only those that are necessary. B Biocompatibility and Safety By definition, a biocompatible material is one that does not elicit a response from the host. As many biomaterials for tissue engineering are biodegradable, it is important to confirm that the degradation particles are also biocompatible in all aspects of this definition. A successful biomaterial should be tested for biocompatibility and safety to ensure that the material is not thrombogenic, cytotoxic, or conducive of an excessive inflammatory or immune response. In certain cases, it may be necessary to demonstrate that the biomaterial or its degradation products are not carcinogenic. The biocompatibility parameters encompass the traditional criteria for how biomaterials have been chosen in the past. New demands for tissue engineering applications require that biomaterials be actively involved and promote a “healthy” response by stimulating the cells, while still avoiding a negative reaction.

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C Biodegradability In cartilage tissue engineering, biodegradable biomaterials are often expected to fill an early biomechanical role as well as facilitate functional cartilage synthesis [49]. Biodegradable biomaterials have the distinct advantage of being able to provide a temporary environment for the cells to grow until they can form their own extracellular matrix that eventually replaces the temporary biodegradable scaffold. The transition from artificial to natural environment should be a rather seamless one in order that the cells do not experience a sudden change in their physical environment (Fig. 3), which could change their state of differentiation. The kinetics of biodegradation of the biomaterial should be controlled precisely to give enough time for the cells to lay down their own extracellular matrix and regenerate the injured cartilage, and at the same time to ensure that the scaffold does not last longer than needed. Biodegradable scaffold materials should also fulfill the biocompatibility or safety requirement discussed previously. D Biological Properties Besides being biocompatible and safe, it is often thought that the biomaterial developed for a tissue engineered construct must attempt to mimic the original tissue on a biological level. As discussed earlier, the extracellular matrix of cartilage is comprised from many collagenous and noncollagenous proteins, proteoglycans, ions, and other small molecules, all of which can direct cellular behavior [25]. Therefore, the creation of successful tissue engineered cartilage may hinge upon the ability to identify, prioritize, and integrate into tissue engineered constructs the molecules that most heavily affect chondrocyte biology and metabolism. The most obvious to focus upon would be those that have direct contact with the chondrocyte, either by occupying membrane receptors or integrins. These molecules may include growth factors, cytokines, or other larger matrix molecules. Growth factors such as the TGF- family, IGF-1, FGF, and PDGF, have several effects on chondrocytes grown in vitro and in cartilage explant studies and have in some cases been shown to enhance the cartilage repair process when injected intra-articularly [50,51]. In general these growth factors are thought to invoke an anabolic response by enhancing collagen and proteoglycan production [52] as well as tissue inhibition of metalloproteinase (TIMP) production [53]. Given this evidence, it may be beneficial that the biomaterial serves as a source of any one or a combination of the different growth factors in order to achieve optimal biological success [54]. Cytokines are also known to play a role in chondrocyte biology and cartilage repair [55]. In general, cytokines are associated with a catabolic response including suppression of -1 (II) procollagen mRNA expression as well as its synthesis [55]. Therefore, because healthy cartilage requires a balance of anabolic and catabolic reactions, it may be necessary to expose the tissue engineered construct to cytokines as well. Recent research has also focused on adding “bioactive” peptides to the biomaterial scaffold in order to mimic the environment to which a chondrocyte would normally be exposed [56]. This is done in order to occupy different integrins that may be present on the cell membrane to trigger intracellular signaling pathways [48]. E Mechanical Properties Prescribing the proper mechanical properties for a biomaterial to be used in a tissue engineering application is critical to the success of the implant. This criterion is of most importance, especially when attempting to repair or replace a tissue such as cartilage whose role

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is mechanical in nature. The end goal is that the biomaterial scaffold be able to sustain normal joint loads to allow the patient to return to pain-free activity. Additionally, the mechanical influence on the cells within a biomaterial will be related to the mechanical properties of the scaffold, the boundary loading conditions acting on the construct, and the number and characteristics of focal contacts of the cell to the scaffold. All of these factors will contribute to the cell’s ability to respond to both mechanical and biologic signals, and subsequently contribute to the ultimate biological success of the implant [57]. The determination of mechanical properties that a biomaterial must possess primarily requires a deep understanding of the mechanical properties of the tissue being replaced, the loads normally experienced by the tissue, and the cells’ response to these applied loads. For articular cartilage, the anisotropic, inhomogeneous, nonlinear, and viscoelastic properties of cartilage seem to be critical to the tissue’s proper function in joint contact and lubrication. Additionally, there is increasing evidence that the regulation of chondrocyte biology in vivo is also influenced by these properties and by the biomechanical interactions between the cell and its extracellular matrix. The mechanical environment of the chondrocyte will be dictated by the properties of the material surrounding the cells as well as the applied loads [46]. However, since it will be nearly impossible to duplicate all of the mechanical properties of cartilage, efforts must focus on discovering which properties take priority over the others on a biological level. This has led to the idea of “functional tissue engineering,” a concept that emphasizes keeping in mind that the tissue being created should be functional in a biomechanical sense as well as functional in a biological and biochemical sense [57]. F Ease of Handling and Sterility After the tissue engineered construct is developed and built, the engineer is faced with the challenge of being able to handle, ship, and finally see implantation of the construct into its final biological environment, the human body. The tissue engineered construct must be easy to handle by both the technicians who produce it as well as the surgeons who implant it so that the implant is not destroyed or contaminated during packing, transit, or surgery. Therefore the biomaterial must be in a form that can be readily handled as well as sterilized. In summary, the aforementioned criteria represent some of the minimal essential requirements for successful tissue engineered repair. To determine whether or not a tissue engineered construct meets these criteria, standards of testing and evaluation in vitro and in vivo, in animal models and in human patients, must still be established and universally accepted. Despite the absence of such standards, several biomaterials have been investigated for applications in cartilage tissue engineering. In the following discussion we will attempt to classify the different types of biomaterials used in cartilage tissue engineering studies and describe their performance under the functional criteria mentioned. The materials reviewed are classified primarily according to their source as synthetic or natural. Synthetic materials are defined as those materials that are fabricated in the lab and that are not present naturally in the mammalian body. Natural materials are modified derivatives of proteins and polymers that may be present in mammalian and nonmammalian tissues. IV SYNTHETIC BIOMATERIALS IN CARTILAGE TISSUE ENGINEERING Various polymers have been investigated for use in tissue repair and reconstruction (e.g., cartilage [58–60], meniscus [61], bone [62], nerve [63], and cardiac tissue [64]), in bone fixation [65–67], in burn wound coverings [68], in surgical sutures [69–71], and in drug de-

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livery [72–74]. In principle, any biodegradable polymer that meets the aforementioned criteria may be used. However, among the most commonly studied synthetic polymers in tissue engineering are the biodegradable linear aliphatic polyesters poly(-hydroxy acids). These include polyglycolic acids (PGA), polylactic acid (PLA), and their derivatives and lactide–glycolide copolymers (PLGA). A -Hydroxy Acid Polyesters PLA, PGA and PLGA copolymers are synthetic, biodegradable -hydroxy esters. PLA polymers are usually prepared from the cyclic esters of lactic acids. Common forms of PLA polymers include the L() form of lactic acid—poly(L-lactic acid or PLLA)—and the mixture of D() and L() lactic acids—poly(DL-lactic acid). PGA is a homopolymer of glycolic acid (hydroxyacetic acid). Glycolic acid is initially reacted with itself to form the cyclic ester glycolide, which in the presence of heat and a catalyst is converted to a high molecular weight linear-chain PGA polymer. PLA and PGA polymers degrade by hydrolysis without toxic degradation products. PLLA degrades at slower rates than PGA. The biodegradability of the matrix in general depends upon the molecular weights of the polymer (PLA, PGA) and the ratio of the monomers in a copolymer (PLGA) (Fig. 4) [75]. The higher molecular weights result in polymer matrices that degrade over longer periods of time; while lower molecular weights result in shorter matrix lives. The detailed physicochemical characteristics and methods of preparing these polymers are beyond the scope of this chapter and can be found in other dedicated references [76]. The discussion here is limited to tissue engineering applications of such polymers and in particular to those dealing with cartilage. Several groups have investigated the use of polylactic acid and its derivatives as scaffold materials to regenerate cartilage tissue. For example, in 1995 Bujia et al. [77] utilized

Figure 4 Poly(lactic acid) composition affects the in vivo half-life of PLGA polymers. (From Ref. 75.)

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bioresorbable polymer fleeces of PLA as temporary cell scaffolds to “establish three-dimensional cultures of human chondrocytes.” To improve cell retention, the polymer surface was coated with poly-L-lysine before cell seeding and the cell–polymer constructs were then encapsulated with agarose and placed in perfusion chambers. The scaffold and culture conditions maintained the differentiated phenotype of the chondrocytes as manifested by synthesis of proteoglycans and collagen type II [77]. Others have investigated the suitability of commercially marketed porous poly(DLlactic acid) polymer (a dental dressing material for large animals, ADD Cube, THM Biomedical, Inc./OSMED, Duluth, MN) as a scaffold material for delivering perichondrocytes into full-thickness articular cartilage defects in the femoral condyles of adult New Zealand white (NZW) rabbits [78,79]. In vitro fluorescent visualization of the live and dead cells in the scaffold using confocal microscopy demonstrated that the perichondrocytes were capable of attaching to and surviving within the porous poly(DL-lactic acid) scaffold [79]. Six weeks after implantation, 96% of the experimental knee defects demonstrated repairs consisting of smooth, firm neocartilage that resembled the color and texture to the host articular cartilage and stained positively for cartilaginous protein-specific antibodies. These results suggest that porous poly(DL-lactic acid) scaffolds have the ability to retain and maintain the viability and phenotype of the chondrocytes in vitro and in vivo and to support the growth of cartilaginous repair tissue in vivo. The study however did not address the temporal changes in the mechanical quality of the cartilaginous repair tissues and the biodegradation kinetics of the implanted scaffold [78]. A later study demonstrated the efficacy of the repair of osteochondral defects in the femoral condyles of NZW rabbits over 12 months using allogenic perichondrocyte–polylactic acid grafts and showed that 85% of the grafted knee defects were filled with what was described as a cartilaginous material [80]. Although, none of the specimens displayed completely normal cartilage histology at 1 year. Biochemically, type II collagen was the dominant constituent of the repair matrix, whereas sulfated glycosaminoglycan content appeared subnormal at all time periods. The “confined-compression creep” modulus of the repair tissue at all time points up to 12 months was not significantly different from that of the “normal” nonoperated contralateral knee cartilage (0.28  0.11 MPa) [80]. More recently, the same group demonstrated that autologous perichondrocyte–polylactic grafts improved the histological appearance of the cartilaginous repair tissue compared with allogenic grafts. This observation implies that the scaffold material is but one factor that may affect the success of the tissue engineered repair, and that cell source may be of equal or greater importance. However, even with the autologous perichondrocytes, the concentration of glycosaminoglycans in the neocartilage matrix remained suboptimal, the surface of the repair tissue was depressed relative to the surrounding host cartilage, the histological appearance was not consistent with normal cartilage features, and the biomechanical properties of the repair tissue were inferior to normal cartilage [81]. Fibrous polyglycolic acid and porous poly(L-lactic acid) scaffolds have also been seeded with cultured bovine and human chondrocytes and evaluated in vitro and in vivo [82]. After 6 weeks in vitro, the chondrocytes grown on PGA increased in number density by more than eight fold, almost doubling the cellular growth rate on PLLA scaffolds. Further, chondrocytes grown on PGA scaffolds produced higher sulfated glycosaminoglycans (GAGs) compared to cells grown on PLLA. When implanted in vivo, the original polymer scaffold shape of both PGA and PLLA was maintained for up to 6 months after implantation. When retrieved after 6 months, the implants of both polymers displayed a hyaline cartilage-like appearance and contained sulfated GAGs and type II collagen [82].

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Proteoglycan and collagen synthesis rates have been compared for chondrocytes cultured on various scaffolds including nonwoven PGA fiber mats (Davis and Geck, Danburry, CT), PLGA flat woven mesh (Vicryl™, Ethicon, Sommerville, NJ), fibrillar bovine type I collagen (American Biomaterials, Plainsboro, NJ), and a nylon mesh weave [83]. Collagen promoted the most efficient cellular attachment, whereas nylon was the least efficient. PGA scaffolds fared better in promoting cellular attachment than did PLGA scaffolds. Cells in the collagen scaffolds appeared to be fibroblastic in morphology, while most cells in the synthetic scaffolds without direct attachment to the polymer fibers assumed spherical chondrocytic morphology. PGA and PLGA scaffolds promoted greater rates of proteoglycan synthesis compared to the nylon and collagen matrices (Fig. 5a). Proteoglycan synthesis was also significantly increased in the PGA scaffold compared to the PLGA matrix. The collagen scaffold stimulated close to a four-fold increase in the synthesis of collagen (Fig. 5b) [83].

Figure 5 Rates of PG and collagen synthesis in different scaffold materials at 5 weeks in a closedloop recirculation culture system. (From Ref. 83.)

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Hybrid scaffolds made from synthetic polymers and low-melting agarose type IX have been utilized for the reconstruction of septal cartilage [84]. Chondrocytes from human septal cartilage were suspended in 1.5% (w/v) agarose IX and seeded onto either Ethisorb, a polyglycolic acid/poly-L-lactic acid (90:10) copolymer (Ethicon, Norderstedt, Germany), or V 7-2, a pure PLLA (Institute for Textile and Process Engineering, Denkendorf, Germany). Both polymer scaffolds were pretreated with poly-L-lysine to improve chondrocyte attachment to the scaffold fibers. The cell-seeded polymer scaffolds were then encapsulated with 4% (w/v) agarose IX. After 1 week in an in vitro perfusion culture chamber, the cell-loaded scaffolds were implanted into nude mice. PGA/PLLA (90:10) scaffolds degraded within 3 weeks and showed no significant differences from normal human septal cartilage in the amount of collagen II at 24 weeks. The pure PLLA scaffolds contained significantly smaller amounts of collagen II and showed minimal degradation even after 24 weeks [84]. These results suggest that the delicate balance between synthesis rates of cartilage proteins and degradation rates of the scaffold material may ultimately determine the extent of tissue regeneration. Other hybrid scaffolds made from fibrin glue seeded with bovine articular chondrocytes and then incorporated into PGA/PLLA (90:10) scaffold meshes (Ethisorb, Norderstedt, Germany) [85] were recently developed. These hybrid scaffolds were then implanted in subcutaneous pockets in the dorsa of nude mice. The implants were retrieved at 6 and 12 weeks and biomechanically evaluated using a mechanical indentation instrument. The mechanical properties of the implants improved significantly between 6 and 12 weeks. By 12 weeks, the compressive Young’s modulus (~15 MPa) and failure stress (~5 MPa) of the implants were almost equal to values of cartilage specimens from human nasal septum, albeit they were only 25–30% of normal bovine articular cartilage [85]. Recent advances in polymer science and three-dimensional tissue culture systems have yielded significant developments, including the introduction of multiphasic polymer scaffolds [86,87] and RGD peptide modified polymers [88]; the pretreatment with factors such as poly-L-lysine, as described previously [77,89,90]; the introduction of injectable polymers that can be crosslinked in situ [91–95]; and the use of bioreactors, perfusion systems, and spinning flasks to promote more efficient distribution of nutrients and uniform seeding of cells [47,96–100]. B Multiphase Polymer Scaffolds Multiphase implants from D,L-PLA-co-PGA with additives such as PGA fibers, 45S5 Bioglass®, and calcium sulfate locally manipulated to control the physicochemical properties of each phase have been described previously [86]. The implants consisted of three phases. The articulating surface phase (100  30 m) consisted of a thin fully dense film composed of 75% D,L-PLA 25% PGA. The porous cartilage phase (1.2 mm) was composed of a 75% D,L-PLA 25% PGA composite, with or without 10% PGA fiber reinforcement (compressive stiffness, Ec  12  1.5 and 32  2.1 MPa, respectively). The subchondral bone phase (2.7 mm) was made from either a 75% D,L-PLA 25% PGA composite (Ec  12  1.5 MPa); a 20% PGA fiber–reinforced 75% D,L-PLA 25% PGA composite (Ec  48  5.4 MPa); a 55% D,L-PLA 45% PGA composite with 20% Bioglass particles (Ec  0.3  0.06 MPa); or a 75% D,L-PLA 25% PGA composite with 50% medical grade calcium sulfate (Ec  1080  484 MPa) [86]. The phases were then glued together and 3-mm cylindrical implants were cored out and loaded with chondrocytes retrieved from a cartilaginous rib biopsy of the experimental animals (goats). Each implant contained 30.49 103  11.41

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103 cells. Implants containing autologous cells were then used to surgically repair osteochondral defects drilled in high- and low-weight-bearing regions of the medial femoral condyle and the medial portion of the patellar groove. Qualitative histological assessment at 16 weeks showed that all implant groups had a high percentage of hyaline cartilage and good bone restoration and neocartilage integration with no apparent differences among the implant types. No implant-related differences in the stiffness of the neocartilage tissue could be observed, with stiffness values averaging 53.4  1.9 and 57.8  2.0 for the condyle and patellar groove sites, respectively (ranging between 73 and 81% of normal site-matching healthy cartilage) [86]. Other biphasic cartilage–bone composites have been recently developed by seeding articular chondrocytes and periosteal cells and polymer scaffolds of varying compositions [87]. To make the cartilage phase, calf articular chondrocytes were dynamically seeded on fibrous, nonwoven PGA meshes (2 mm thick, 97% porosity) (Albany International Research Company, Mansfield, MA). The bone phase was made by dynamically seeding periosteal cells on foams of 80:20 poly(lactic-co-glycolic acid) and polyethylene glycol (PLGA/PEG) (1.5 mm thick, 85% porosity). PGA–chondrocyte constructs were cultured in chondrogenic conditions, whereas the PLGA/PEG–periosteal cell constructs were cultured in osteogenic conditions independently for 1 or 4 weeks. The two phases were then sutured together and cultured for an additional 4 weeks in the osteogenic conditions (Fig. 6). The biphasic constructs displayed distinct well-defined regional histology that resembled cartilage and bone tissue. The sulfated GAG content (assessed by image analysis of safranin-O staining) in the cartilaginous phase increased significantly with time before and after joint culture, whereas the mineralization in the bone phase (assessed by image analysis of immunohistochemical staining against osteocalcin) was not measurable until 4 weeks of independent culture and remained constant thereafter through 4 weeks in joint culture

Figure 6 Histology of constructs and osteochondral composites: (A) 1-week cartilage construct (safranin-O staining); (B) 1-week bone construct; (C) composite of (A) and (B) after 4 weeks of combined culture; (D) 4-week cartilage construct; (E) 4-week bone construct; and (F) composite of (D) and (E) after 4 weeks of combined culture. (From Ref. 87.)

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[87]. These results demonstrate that multiphasic implants can be assembled from heterogeneous populations of cells and scaffolds of varying physicochemical characteristics and emphasize the importance of understanding and controlling culture conditions for optimal and controlled tissue regeneration. C Peptide-Modified Polymers Modifying polymer scaffolds to integrate biosynthetic ligands offers the potential to “engineer” tissues by eliciting controlled cellular responses. In 1997 Cook et al., synthesized and characterized arginine-glycine-aspartic acid (RGD) peptide modified poly(lactic acidco-lysine) (PLAL) [88]. Lysine containing monomers were copolymerized with L-L-lactide to produce the PLAL polymer. The synthetic peptides were then covalently attached to the PLAL that was later blended with poly(lactic acid) (PLA) in varying proportions ranging between 10:90 and 50:50 PLAL/PLA [88]. Peptide coupling efficiency, defined as the ratio of peptide to lysine, was determined by measuring primary amino acid group concentrations using colorimetric assays and x-ray photoelectron spectroscopy (XPS). Peptide coupling efficiency ranged from 1 to 26%. When the scaffolds were seeded with bovine aortic endothelial (BAE) cells, the cells’ mean spread area increased by six fold after a 4-h incubation period on RGD-PLAL surfaces and was nearly four times greater than the mean spread area of cells grown on PLAL or PLA surfaces without RGD peptides [88]. These results demonstrate that peptide-modified biopolymers may in fact boost the response of cells. This may ultimately have significant implications in tissue engineering once in vivo and in vitro studies are conducted to better understand how to capitalize on this concept. D Injectable Polymers Injectable polymers that can be crosslinked in situ offer a major advantage over conventional polymer scaffolds that would require surgical installation [49,91,92,94,95]. Injectable poly(ethylene oxide) gels, which are biocompatible and biodegradable synthetic polymers, have been utilized for the encapsulation of calf articular chondrocytes and have been shown to maintain cellular viability and phenotype [95]. PEO–chondrocyte gels with cellular densities of 10 10 6 per milliliter were injected subcutaneously in nude mice. DNA and sulfated GAG analysis of the injected PEO–cell suspensions at 6 and 12 weeks postimplantation in vivo confirmed the proliferative and synthetic activity of the chondrocytes [95]. In 1999 Elisseeff et al. investigated the use of injectable PEO-based hydrogels that can be transdermally photopolymerized in situ following injection or implantation [91,92]. PEO–dimethacrylate (PEODM) and PEO hydrogels (10:90, 20:80, 30:70, and 40:60 PEODMPEO) seeded with bovine articular chondrocytes were injected subcutaneously in athymic mice. After injection, the hydrogels were transdermally crosslinked using ultraviolet (UV) light at an intensity of 2 mW/cm 2 for 3 min. Postinjection analysis of the retrieved implants after 2 to 7 weeks demonstrated that the chondrocytes survived the photopolymerization protocol. DNA content remained almost constant at all time points and was not significantly affected by the percentage of PEODM in the hydrogel. Sulfated GAG content decreased significantly with increasing ratio of PEODM, ranging from 2.8% per wet weight for the 10% PEODM constructs to 1.5% per wet weight for the 40% PEODM constructs. Collagen content was not affected by the PEODM ratio but was significantly increased with time, reaching values of 6.5% per wet weight at 7 weeks of which 35% was type II collagen [92]. Such injectable polymers, which can be polymerized in situ via transdermal UV light exposure or via some other means including temperature and pH

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changes, offer minimally invasive methods for tissue engineering. More work is still needed to determine the efficacy of in situ injection procedures in reconstructing osteochondral lesions of irregular geometries in animal and human models. V NATURAL BIOPOLYMERS Collagen; derivatives of various naturally occurring polysaccharides, such as glycosaminoglycans and hyaluronic acids; as well as other biomaterials extracted from sea creatures, such as alginic acid, agarose, and the GAG analog chitosan are typically categorized as natural biopolymers. The biological origin of these natural biopolymers makes them an attractive choice for tissue engineering considerations. However, the disparity in their mechanical and biological functions and properties warrants thorough understanding of their characteristics. Therefore, these natural biopolymers have been routinely studied as potential scaffold materials for tissue engineering applications. A Collagen Collagen is the most abundant protein in the body, constituting the fundamental structural protein of virtually every tissue, including skin, tendon, ligament, bone, cartilage, and blood vessels. There are at least 17 distinct polypeptide chains that combine to form more than 10 variants of collagen [101]. The extracellular matrix of articular cartilage contains collagen types II, VI, IX, X, and XI. Collagen type II is almost exclusively synthesized by chondrocytes and accounts for nearly 80% of the collagen in articular cartilage, while the remaining 20% is mostly comprised of types IX and XI [102]. The fibrillar forms of collagen, types II, IX, and XI, form an interweaving mesh of fibrils that is responsible for the tensile properties of the cartilage extracellular matrix [49]. Collagen type VI is localized in the pericellular matrix of chondrocytes in beaded filament form [102]. Collagen type X is almost exclusively synthesized by hypertrophic chondrocytes and is thought to make up the hexagonal filamentous network of fibrils observed in hypertrophic cartilage [102]. While it may seem logical that collagen type II should be the material of choice in cartilage tissue engineering, very few studies have indeed used collagen II in such endeavors. Instead, most investigators used scaffolds made from type I collagen. This is mainly because collagen type I is abundantly available from a great number of tissue sources and can be cost-effectively isolated for commercial, therapeutic, and investigative uses. Collagen I is easily isolated and purified from collagen-rich tissues such as bovine tendon into soluble molecules, using enzymatic digestion with proteolytic enzymes such as pepsin, or into fibrillar forms using sequential salt extraction and enzymatic digestion of noncollagenous proteins and lipids [103]. Advances in the biotechnology of collagen have enabled the fabrication of different collagen matrices including collagen gels, collagen solutions, porous collagen sponges, and filamentous and tubular collagen membranes for various medical and investigative applications [103]. Such biodegradable matrices have demonstrated natural biocompatibility due to the homology of collagen molecules from bovine and human tissues. In fact, the in vivo degradation of bovine collagen follows enzymatic pathways similar to tissue remodeling during healing [103]. The rate of in vivo degradation and the mechanical properties of resorbable collagen scaffolds depend on the porosity and the extent of crosslinking of the scaffold [103]. The main disadvantage of collagen as a scaffold for cartilage tissue engineering, however, is perhaps its inability to preserve the chondrocytic phenotype of cells.

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Cells seeded on collagen scaffolds have been shown to display fibroblastic morphology and to effect collagen shrinkage [104,105]. Researchers have investigated the use of various forms of collagen as potential scaffolds for articular cartilage tissue engineering [16,106–108]. For example, marrow and periosteal mesenchymal progenitor and cells have been suspended in type I collagen gel scaffolds that were transplanted into osteochondral defects in rabbit femoral condyle [16]. The implanted marrow and periosteal cells exhibited similar patterns of differentiation into the chondrocytic phenotype initially producing a fairly thick superficial layer of a cartilage matrix that significantly stained with toluidine blue. Cancellous bone filled the remainder of the defect underneath the cartilage layer. But by 24 weeks, the superficial layer became very thin and did not stain for toluidine blue, and the repair tissue surface was irregular, while the subchondral bone appeared to have been completely regenerated. The cell/collagen repair tissues were three-fold more compliant (i.e., less stiff) than normal articular cartilage [16]. Others have tried to modulate type I collagen matrices by covalently binding GAGs including chondroitin sulfate (CS) to the collagen fibrils using 1-ethyl-3-(3-dimethyl aminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS) [109]. The ultrastructural localization of CS in the GAG-augmented collagen matrices showed discrete fil-

Figure 7 Ultrastructural localization of GAGs in (A) CS-augmented collagen matrices and in (B) bovine tracheal cartilage. Arrows indicate individual GAGs. Arrowheads indicate individual proteoglycans. (From Ref. 109.)

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Figure 8 Cell morphology percentages of canine chondrocytes in type I and type II collagen matrices at 3 h (3h), 7 days (7d), and 14 days (14d). (From Ref. 104.)

aments of single CS molecules throughout the matrix in a pattern resembling the localization of GAGs in the pericellular matrix of chondrocytes in bovine tracheal cartilage (Fig. 7) [109]. When implanted in vivo, these GAG-augmented collagen matrices invoked a mild inflammatory response that decreased with time and retained their scaffold integrity and porous structure over 10 weeks [110]. Such GAG-augmented collagen scaffolds, however, have not been tested to determine whether or not they possess the mechanical and biochemical characteristics necessary to provide adequate support for the cells with subsequent growth and matrix synthesis. The ability to maintain chondrocyte morphology and phenotype has been investigated and compared in vitro for two porous scaffolds: a chondroitin-6-sulfate–augmented collagen I scaffold (bovine) and a collagen II scaffold (porcine) in vitro [104,105]. The GAG–collagen I scaffolds showed a 30% contraction in the 4-mm cylindrical scaffolds after 7 days, while the collagen II scaffolds showed almost no shrinkage. After only 3 h in culture, almost 70% of the chondrocytes seeded on the GAG–collagen I scaffolds displayed an elongated, fibroblastic morphology without significant change over 14 days in culture (Fig. 8). By contrast, nearly 70% of the chondrocytes in the collagen II scaffolds displayed spherical morphology with some pericellular matrix formation that stained positively with safranin-O and against collagen II–specific antibodies [104]. The GAG content increased in both scaffolds with time, but at 14 days the collagen II scaffolds contained almost double the amount of GAGs in the collagen I scaffolds [104]. When implanted in vivo to repair partial thickness chondral defects in canine trochlea, the autologous chondrocyteseeded collagen II scaffolds demonstrated the greatest amount of reparative tissue that

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comprised mixed types of mostly fibrous tissue and very little hyaline cartilage-like tissue [108]. The chondrocyte-seeded GAG–collagen I scaffolds resulted in incomplete filling of the defects with mostly fibrous tissue and very little hyaline cartilage-like tissue. Both scaffolds stained positively for collagen II and safranin-O, although the staining intensity in the collagen II scaffolds was more intense [108]. These results demonstrate that collagen II has some advantage over collagen I as a potential scaffold material for cartilage tissue engineering applications. More studies must be completed to evaluate other functional characteristics of such scaffolds. B Hyaluronic Acid Hyaluronic acid (HA) is a naturally occurring polysaccharide consisting of alternating Dglucuronic acid and N-acetyl D-glucosamine monomers. It is abundant in the mesenchyme of developing embryos [111]. It is also present in the extracellular matrix of articular cartilage and in other connective tissues as well as in the synovial fluid. In articular cartilage, HA plays a vital role in the assembly of proteoglycan molecules to form aggregating proteoglycans [49]. Therefore, HA plays an important role in the maintenance of cartilage hydration and structural integrity. Hyaluronic acid also plays important roles in tissue repair by promoting mesenchymal and epithelial cell migration and differentiation, improving angiogenesis, and enhancing collagen production [112,113]. Because of its important biological roles, HA has been used in therapeutic products including a product for intraocular injection during cataract removal, corneal transplant and glaucoma surgery (Amvisc®, Anika Therapeutics, Woburn, MA), injectable agents to relieve osteoarthritic pain of the knee (Synvisc®, Biomatrix, Inc., Ridgefield, NJ; Orthovisc®, Anika Therapeutics, Woburn, MA), and a dressing for severe burn wounds (Hyaff-11, Fidia Advanced Biopolymers, Padua, Italy). Chemically modified derivatives of HA have been recently developed to tailor the physicochemical properties of the material while maintaining its biological properties. For example, Hyaff-11 was synthesized through the complete esterification of carboxyl groups of the glucuronic acid with benzyl groups. The biocompatibility of Hyaff-11 appears to be quite good and the material degrades slowly within 60 days in vitro and 110 days in vivo synchronous with neotissue formation [114,115]. It should be noted, however, that the rate of degradation depends upon the degree of esterification. While completely esterified HA scaffolds such as Hyaff-11 have half-lives of several months, partially esterified derivatives of HA scaffolds have been shown to degrade in weeks. The material can also be sterilized effectively with -irradiation [114]. The use of the HA derivative Hyaff-11 as a scaffold for articular cartilage tissue engineering has been explored [111,114,116,117]. For example, avian chondrocytes were shown to have easily adhered and proliferated onto nonwoven meshes of the HA-based scaffolds (Hyaff-7 and Hyaff-11) and synthesized a glycosaminoglycan- and collagen II–rich extracellular matrix [114]. Scaffolds made from Hyaff-11 and a crosslinked derivative of HA that is generated by condensation (ACP) were loaded with rabbit mesenchymal progenitor cells and compared against a control of porous, fibronectin-coated calcium phosphate ceramic cubes (Fig. 9). Hyaff-11 scaffolds retained 90% more cells than did the ACP scaffolds and the fibronectin-coated ceramics. When implanted subcutaneously in nude mice, Hyaff-11 scaffolds retained their integrity after 3 and 6 weeks, whereas the ACP scaffolds rapidly and completely degraded within 7 to 10 days. Bone, cartilage, and fibrous

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Figure 9 Scanning electron micrograph images of (a) the fibronectin-coated ceramic cube and (b) the Hyaff-11 scaffolds. High magnification SEM of (c) the cell-seeded fibronectin-coated ceramic cube and (d) the cell-seeded Hyaff-11 scaffolds after 2 h in culture. Histological evidence of chondrogenesis in (e) the fibronectin-coated ceramic cubes and (f) Hyaff-11 scaffolds 3 weeks after implantation. (From Ref. 111.)

tissue formation was evident in the pores of the implanted Hyaff-11 and ceramic scaffolds, with the former showing a 30% increase in the relative amount of bone and cartilage per unit area [111]. More recently, Hyaff-11 scaffolds loaded with autologous, bone marrow–derived mesenchymal progenitor cells were implanted into osteochondral lesions in rabbit knees. The implanted scaffolds did not seem to elicit a severe inflammatory response and were completely degraded within 4 months after implantation. Blind histological scoring of the repairs demonstrated that lesions filled with cell-loaded Hyaff-11 scaffolds healed somewhat better than unfilled lesions and those filled with acellular Hyaff-11 scaffolds [117]. Results from these studies demonstrate that the biocompatible derivatives of HA provide adequate support for chondrogenesis, retain cells effectively, and degrade at slow enough rates to protect and promote neotissue synthesis. Data on the mechanical properties of the HA-based repair tissue are still lacking, however.

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C Chitosan Chitosan, a deacetylated derivative of chitin, is a GAG analog that has been gaining recognition in the tissue engineering community as a potential scaffold biomaterial. Chitin and chitosan are the main structural components of the cuticles of aquatic arthropods and insects [118]. Chitosan is a semicrystalline linear polysaccharide of (1→4)-linked D-glucosamine residues with randomly located N-acetyle-glucosamine groups that displays remarkable structural homology with various GAGs and HA [119]. The solubility of chitosan is pH dependent, as it is normally insoluble in aqueous solutions above pH 7.0 and is completely soluble below pH 5.0 [119]. Because of this property, viscous solutions of chitosan can be extruded and then gelled by a simple shift in pH to form gel fibers that can be subsequently drawn and dried [119]. Chitosan is biocompatible and biodegradable [118]. The mechanism of chitosan degradation in vivo is through enzymatic hydrolysis primarily by lysozyme [119]. The rate of degradation is inversely related to the degree of deacetylation or crystallinity. The main degradation products of chitosan are variable-length chitosan oligosaccharides, which were found to stimulate a minimal foreign body reaction in vivo [119]. Perhaps the most intriguing property of chitosan is the ease and control with which porous structures can be created. In 2000 Suh et al. reported that porous chitosan scaffolds were made using a two-step process involving freezing and then lyophilizing chitosan acetic acid solutions [119]. Freezing parameters, including the freezing rate and geometry of thermal gradients, can be varied to control pore size and orientation, respectively (Fig. 10) [119]. By controlling pore size and orientation, the mechanical properties of such scaffolds may be altered. For example, it has been reported that the elastic modulus of nonporous membranes of chitosan (5–7 MPa) was more than 14-fold greater than the elastic modulus of porous membranes (0.1–0.5 MPa) [119]. The tensile strength of the porous chitosan structures is in the range of 30–60 KPa [119]. The compressive properties of chitosan scaffolds have not been reported. Such three-dimensional porous structures may be used in tissue engineering if they can demonstrate biological and structural functionality. Human osteoblasts and articular chondrocytes have been plated onto monolayer films of chitosan (4% w/v in 2% acetic acid) and demonstrated that chitosan promoted viability and maintenance of phenotype, as evident by the spherical shape of these cells as compared to spindle-shaped cells directly plated on noncoated coverslips [120]. RT-PCR analysis and immunostaining revealed that osteoblasts grown on chitosan monolayers expressed collagen type I, whereas chondrocytes grown in similar conditions expressed collagen type II and aggrecan after just 7 days in culture [120]. Another important property of the cationic chitosan is its ability to compound with anionic polysaccharides such as GAGs, resulting in pH-dependent hydrogels [119]. Sechriest et al. seeded bovine articular chondrocytes onto chondroitin sulfate-A (CSA)–augmented chitosan hydrogel monolayers and noncoated polystyrene surfaces [121]. Chondrocytes seeded on polystyrene changed in morphology to become fibroblastic and underwent a significant degree of proliferation. By contrast, chondrocytes seeded on the CSA-augmented chitosan monolayers remained spherical and demonstrated limited proliferation. The chondrocytes were not embedded in the hydrogel monolayer, but rather formed discrete focal adhesion points. Further, chondrocytes seeded on the CSA-augmented chitosan monolayers predominately synthesized collagen II (58%) and proteoglycans [121]. These early reports demonstrate the potential of chitosan as a scaffold bioma-

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Figure 10 Effects of freezing rates and orientation of thermal gradients during freeze-drying on the porosity and pore orientation of chitosan. High magnification SEM of a chondrocyte adhered to chitosan film. (From Ref. 119.)

terial for articular cartilage tissue engineering, although more studies are still needed to biologically and biomechanically characterize the repair response of articular cartilage lesions grafted with chitosan-based scaffolds. D Alginate Alginate, a natural material derived from sea algae (seaweed), is a linear polysaccharide of (1 → 4)-linked -L-guluronic acid (G) and -D-mannuronic acid (M) residues along a linear chain [122]. Divalent cations, such as Ca2, Ba2, and Mg2, have the ability to bind between the guluronic acid residues of adjacent alginate chains in aqueous solutions to form hydrogels [123,124]. Sodium alginate crosslinked with Ca 2 has been used extensively in cell culture and cartilage tissue engineering because of its ability to immobilize chondrocytes while maintaining their spherical morphology and phenotype [122–129].

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Several properties of alginate gels are dependent on the ratio of mannuronic acid to guluronic acid (M/G) [130]. Higher M/G ratios result in smaller pore size gels that are effective in cell entrapment. The M/G ratio is also thought to affect the biocompatibility of alginate preparations. Mannuronic acid–rich alginate preparations have been shown to stimulate macrophages and lymphocytes in vitro [130]. Moreover, commercially available research-grade alginate preparations have been shown to contain various bioactive immunostimulating constituents [130]. Thus, the biocompatibility of alginate has been debated in the literature [49,130,131]. This has led to the development of techniques to obtain highly purified mannuronic acid–rich (68% mannuronic acid residues) alginate gels that showed no mitogenic activity toward lymphocytes in vitro and no mitogen-induced foreign body reaction in vivo [130]. In many cases, the stability of the alginate hydrogel has also been unpredictable. The strength of calcium- or barium-crosslinked alginate gel was reported to have decreased over time in culture before reaching an equilibrium plateau, possibly due to an outward migration of crosslinking ions into the surrounding medium, regardless of the presence or absence of cells [132]. And contrary to common belief, the strength of calcium-crosslinked alginate gel was not different from the strength of the barium-crosslinked alginate gel [132]. The physical and mechanical properties of alginates have been thoroughly studied [123,133]. The elastic compressive and shear moduli of sodium alginate gels increased significantly with an increase in the concentration of the alginic acid. For example, increasing the concentration of sodium alginate gel from 1 to 3% (w/v) resulted in nearly a ninefold increase in the equilibrium compressive modulus (from nearly 0.9 to 8 KPa, respectively), and a tenfold increase in the equilibrium shear modulus (from nearly 0.35 to 3.5 KPa, respectively) [123]. On the other hand, physiological levels of Na decreased the equilibrium compressive and shear moduli of 2% (w/v) alginate hydrogels [123]. This latter observation is rather important as it explains one of the likely degradation mechanisms of calciumcrosslinked alginate gels in vivo. Various researchers have investigated the ability of alginate gels to act as a scaffold material for chondrocytes to regenerate cartilage tissue [128,134]. In 1996 Paige et al. demonstrated that when implanted in vivo, calcium-crosslinked alginate discs seeded with chondrocytes formed hyaline-like cartilage regardless of Ca2 concentration (15 to 100 mM) or alginate concentration (0.5 to 4.0%w/v) used in gel crosslinking. However, increasing the cell seeding density (0 to 10 million cells/mL) significantly increased the strength of the implanted alginate gels over time, suggesting perhaps a correlation between cell density, matrix synthesis, and biomechanical properties [128]. Others have injected allogenic chondrocyte-seeded alginate solutions into full-thickness articular cartilage defects in the rabbit and crosslinked the alginate in situ using Ca2 (50 mM). Unlike controls filled with acellular alginate gels that generated a fibrous meshwork containing infiltrating fibroblast-like cells, the chondrocyte-seeded alginate gels generated cartilage-like tissue that stained for GAGs [134]. Because of the inability of alginate to specifically bind to receptors on cell membranes, the cells appear to be floating in the pores of the gel while maintaining a spherical morphology that is thought to be crucial to the maintenance or even the inception of a differentiated chondrogenic phenotype. This feature has proven advantageous for novel cartilage tissue engineering approaches using undifferentiated stem cells. Recent studies have demonstrated that the suspension of adipose tissue–derived stromal cells in alginate beads under controlled culture conditions ultimately led the cells to chondrogenic differentiation.

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This was manifested by the increased synthesis of cartilage associated proteins such as aggrecan, collagen types II and VI, and link protein [135]. While these studies demonstrate that alginate gels promote chondrogenesis, the success of tissue engineering applications using alginates is hindered by the poor mechanical properties and handling characteristics of alginates as a scaffold material. E Agarose Agarose is a purified extract from sea creatures such as agar or agar-bearing algae. Structurally, agarose is a linear polymer that consists of alternating -D-galactose and 3,6-anhydro--L-galactose units [136]. Agarose has many biological and biomedical applications especially as a research tool. For example, agarose is primarily used in electrophoresis to separate nucleic acids in northern and southern blots [137,138]. Agarose is also used in the form of beaded or crosslinked gel matrix in chromatographic separation of DNA containing interstrand crosslinks such as those produced by many cancer chemotherapeutic drugs [139]. In addition, agarose has been utilized in tissue culture and tissue engineering applications [140–145]. Agarose solutions exhibit hysteresis in the liquid-to-gel and vice versa transitions [136]. Therefore, liquid agarose solutions can be prepared at temperatures higher than the agarose melting point and can then be gelled and hardened by a decrease in temperature to the gelling point. The gelling temperature is dependent on the constituents and concentration of agarose. Low–melting temperature agarose that can be made into aqueous solutions at physiological temperatures (~37°C) and then gelled at 0 to 4°C is most suitable for mixing with cell suspensions without compromising cell viability [141]. The mechanical properties of the agarose gels are also concentration dependent. For example, the strength of low–gelling temperature agarose type IX gels, defined as the force required for fracturing the gel, varies between as low as 10 KPa at 1.0%w/v to greater than 40 KPa at 1.5%w/v [146]. Increasing agarose concentration also increases the confined compression equilibrium modulus. For example, 5%w/v agarose gel constructs are nearly 2.6 times stiffer than 3 and 13 times stiffer than 1%w/v agarose gel constructs (~65 KPa, 25 KPa, and 5 KPa, respectively) (Fig. 11) [147]. It is also noteworthy

Figure 11 Equilibrium aggregate moduli of alginate and agarose constructs of varying compositions. (From Ref. 147.)

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that the agarose gels are significantly stiffer than similar alginate constructs having the same w/v composition. For example, the 5% agarose gel constructs are almost 4.3 times stiffer than the 5% alginate gel constructs (Fig. 11) [147]. The viscoelastic and dynamic mechanical properties of calf chondrocyte–seeded agarose gels (~1–2 107 cells/mL in 2%w/v agarose solution) have been determined over 70 days in culture [141]. The confined compression equilibrium modulus of the cell-seeded gels increased from 13.1 KPa at day 1 to 103  7 KPa (~22% of normal calf cartilage) by day 35 and then to nearly 150 KPa (~32% of normal calf cartilage) by day 70 [141]. The acellular gels maintained their day 1 values without any significant change up to 35 days. The hydraulic permeability decreased from 40 1015 m4N1s1 at day 1 to 4.15 1015 m4N1s1 (~145% of normal calf cartilage) at days 35 and 70 [141]. These changes in stiffness and permeability were due to the synthesis of a cartilaginous matrix as manifested by an increase in GAG concentration within the gel over time. At day 36, the total DNA content, total GAG content, GAG release rate, and GAG accumulation rate in the chondrocyteseeded agarose gel plugs were nearly 23, 28, 29, and 15% of normal calf cartilage explants, respectively [141]. These results demonstrate that agarose gels retain encapsulated chondrocytes effectively, maintain cellular viability, and promote synthesis of a mechanically functional cartilage-like matrix. Agarose gels have also been used as a model to study chondrocyte biosynthetic response to mechanical compression in vitro [140,142,143,148–150]. For example, Buschmann in 1995 demonstrated that increasing static compression (0–75% relative to free swelling controls) of chondrocyte-seeded agarose gels (2 107 cells/mL in 3%w/v agarose solution) decreased [35S] sulfate and [3H] incorporation after 41 days in culture by nearly 45% of uncompressed control levels [140]. By contrast, dynamic compression (30 m at 0.001–1.0 Hz superimposed over a 27% static compression) stimulated anabolic changes manifested by 15–25% and 10–35% increases in [35S] sulfate and [3H] incorporation, respectively, relative to static controls (27% compression) [140]. Others have demonstrated that dynamically loaded chondrocyte-seeded agarose gels (4 106 cells/mL in 3%w/v agarose) exhibit frequency-dependent and spatial variability in the metabolism of GAG and collagen within the gel [143]. And recently dynamic loading was utilized as a physical stimulant of biosynthesis in tissue engineered chondrocyte/agarose gel constructs [147]. Allogenic chondrocyte-seeded agarose gel implants (1.5 10 6 cells/mL in 1%w/v agarose) have been recently used to repair osteochondral defects in rabbit knees (3 mm diameter 1.5 mm depth) [144]. Prior to transplantation, the implants were cultured in vitro at physiologic conditions for 10–14 days during which the cell density increased to 2.5 10 6 cells/mL. The repair tissues were scored histologically based on the intensity and extent of proteoglycan and type II collagen immunostaining, structural features of the various cartilaginous zone, integration with the host cartilage, and morphological features and arrangement of chondrocytic cells [144]. Despite the variability in the results, the allogenic chondrocyte/agarose–grafted repairs had higher semiquantitative scores than control grafts (acellular agarose) [144]. No data were reported on the biomechanical properties of the tissue engineered repair tissues. These studies demonstrate that agarose gels promote cell retention, proliferation, and chondrogenesis in vivo and in vitro. Data on the in vivo mechanical properties, biocompatibility and toxicity, and the balance between degradation and synthesis kinetics of agarose-based tissue engineered cartilage repair are scarce, however.

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VI CONCLUSION The field of tissue engineering offers exciting new possibilities to improve repair or replacement of various tissues and organs. While the re-creation of events involved in true “regeneration” of the native tissue remains elusive at this time, the prospect of function restoration using tissue engineered grafts has nearly been attained. Whether cells are actively utilized in the tissue engineering process ex vivo or passively recruited from the host to aid the repair in vivo, it is quite apparent that the material scaffold of choice influences the repair and remodeling process before and after transplantation. Synthetic polymers can be tailored to have properties closely matching the tissue they are intended to replace. Their degradation rates and byproducts can be carefully controlled, and their ability to attract and support the growth and biological functions of the cells can be enhanced by exploiting our increasing understanding of cell molecular biology. Natural biopolymers have distinct advantages stemming from the fact that they play important biological roles in vivo and therefore can be manipulated to play similar roles in tissue engineered grafts. These natural biopolymers are in general not as well characterized as synthetic polymers. While the search for optimal material scaffolds continues, it is also apparent that uniform standards of testing and characterization of these scaffolds need to be developed and adhered to in order to make intelligent use of the information generated from such research. ACKNOWLEDGMENTS Supported in part by the National Institutes of Health grants AR43876, AG15768, and GM08555. REFERENCES 1. Gilbert J. E. 1998. Current treatment options for the restoration of articular cartilage. Am. J. Knee Surg. 11(1):42–46. 2. Menche D. S., Frenkel S. R., Blair B., Watnik N. F., Toolan B. C., Yaghoubian R. S., Pitman M. I. 1996. A comparison of abrasion burr arthroplasty and subchondral drilling in the treatment of full-thickness cartilage lesions in the rabbit. Arthroscopy 12(3):280–286. 3. Passler H. H. 2000. Microfracture for treatment of cartilage detects. Zentralblatt fur Chirurgie 125(6):500–504. 4. Minas T., Nehrer S. 1997. Current concepts in the treatment of articular cartilage defects. Orthopedics 20(6):525–538. 5. Ritsila V., Santavirta S., Alhopuro S., Poussa M., Jaroma H., Rubak J. M., Eskola A., Hoikka V., Snellman O., Osterman K. 1994. Periosteal and perichondral grafting in reconstructive surgery. Clin. Orthop. Rel. Res. 302:259–265. 6. Bugbee W. D., Convery F. R. 1999. Osteochondral allograft transplantation. Clinics in Sports Medicine 18(1):67–75. 7. Bobic V. 1996. Arthroscopic osteochondral autograft transplantation in anterior cruciate ligament reconstruction: a preliminary clinical study. Knee Surgery, Sports Traumatology, Arthroscopy. 3(4):262–264. 8. Sasaki T., Yagi T., Tsuge H., Aoki Y. 1985. Clinical results of osteochondral allografts in the treatment of osteochondral defects. Nippon Seikeigeka Gakkai Zasshi (Journal of the Japanese Orthopaedic Association) 59(12):1073–1088. 9. Brittberg M., Lindahl A., Nilsson A., Ohlsson C., Isaksson O., Peterson L. 1994. Treatment of deep cartilage defects in the knee with autologous chondrocyte transplantation. N. Engl. J. Med. 331(14):889–895.

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131. Turner T. D., Spyratou O., Schmidt R. J. 1989. Biocompatibility of wound management products: standardization of and determination of cell growth rate in L929 fibroblast cultures. J. Pharm. Pharmacol. 41(11):775–780. 132. Shoichet M. S., Li R. H., White M. L., Winn S. R. 1996. Stability of hydrogels used in cell encapsulation: an in vitro comparison of alginate and agarose. Biotechnol. Bioeng. 50(4):374–381. 133. Martinsen A., Skjak-Braek G., Smidsrod O. 1989. Alginate as immobilization material: I. Correlation between chemical and physical properties of alginate gel beads. Biotechnol. Bioeng. 33(1):79–89. 134. Fragonas E., Valente M., Pozzi-Mucelli M., Toffanin R., Rizzo R., Silvestri F., Vittur F. 2000. Articular cartilage repair in rabbits by using suspensions of allogenic chondrocytes in alginate. Biomaterials 21(8):795–801. 135. Erickson G., Franklin D., Gimble J., Guilak F. 2001. Adipose tissue–derived stromal cells display a chondrogenic phenotype in culture. 47th Annual Meeting of the Orthopaedic Research Society, San Francisco, CA (in press). 136. Uludag H., De Vos P., Tresco P. A. 2000. Technology of mammalian cell encapsulation. Adv. Drug Del. Rev. 42(1–2):29–64. 137. Ordovas J. M. 1998. Separation of small-size DNA fragments using agarose gel electrophoresis. Methods Molec. Biol. 110:35–42. 138. Upcroft P., Upcroft J. A. 1993. Comparison of properties of agarose for electrophoresis of DNA. J. Chromatogr. 618(1–2):79–93. 139. Hartley J. A., Souhami R. L., Berardini M. D. 1993. Electrophoretic and chromatographic separation methods used to reveal interstrand crosslinking of nucleic acids. J. Chromatogr. 618(1–2):277–288. 140. Buschmann M. D., Gluzband Y. A., Grodzinsky A. J., Hunziker E. B. 1995. Mechanical compression modulates matrix biosynthesis in chondrocyte/agarose culture. J. Cell Sci. 108(Pt. 4):1497–1508. 141. Buschmann M. D., Gluzband Y. A., Grodzinsky A. J., Kimura J. H., Hunziker E. B. 1992. Chondrocytes in agarose culture synthesize a mechanically functional extracellular matrix. J. Orthop. Res. 10(6):745–758. 142. Elder S. H., Kimura J. H., Soslowsky L. J., Lavagnino M., Goldstein S. A. 2000. Effect of compressive loading on chondrocyte differentiation in agarose cultures of chick limb bud cells. J. Orthop. Res. 18(1):78–86. 143. Lee D. A., Noguchi T., Knight M. M., O’Donnell L., Bentley G., Bader D. L. 1998. Response of chondrocyte subpopulations cultured within unloaded and loaded agarose. J. Orthop. Res. 16(6):726–733. 144. Rahfoth B., Weisser J., Sternkopf F., Aigner T., von der Mark K., Brauer R. 1998. Transplantation of allograft chondrocytes embedded in agarose gel into cartilage defects of rabbits. Osteoarthritis & Cartilage 6(1):50–65. 145. Lee D. A., Frean S. P., Lees P., Bader D. L. 1998. Dynamic mechanical compression influences nitric oxide production by articular chondrocytes seeded in agarose. Biochem. Biophys. Res. Com. 251(2):580–585. 146. Sigma. 1996. Agarose (CAS number 9012-36-6)—product information sheet. Sigma: St. Louis, MO. 147. Mauck R. L., Soltz M. A., Wang C. B., Wong D. D., Chao P.-H. G., Valhmu W. B., Hung C. T., Ateshian G. A. 2000. Functional tissue engineering of articular cartilage through dynamic loading of chondrocyte-seeded agarose gels. J. Biomech. Eng. 122(3):252–260. 148. Lee D. A., Noguchi T., Frean S. P., Lees P., Bader D. L. 2000. The influence of mechanical loading on isolated chondrocytes seeded in agarose constructs. Biorheology 37(1–2):149– 161. 149. Freeman P. M., Natarajan R. N., Kimura J. H., Andriacchi T. P. 1994. Chondrocyte cells respond mechanically to compressive loads. J. Orthop. Res. 12(3):311–320.

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15 Polymeric Biodegradable Hard Tissue Engineering Applications Gamze Torun Köse and Vasif Hasirci Middle East Technical University, Ankara, Turkey

I BONE TISSUE ENGINEERING Bone tissue engineering is designed to mimic the natural process of bone repair. Bone repair is an attractive and natural target for tissue engineering. Tissue engineering of bone requires three important elements. These are cellular components, extracellular matrix scaffolds, and growth and differentiation factors. Cells can be either obtained from an exogenous source or they can be recruited from the local environment. A scaffold matrix must provide a substrate for cellular attachment, proliferation, and differentiation. Growth and differentiation factors guide the appropriate development of the cellular components. A Bone Bone is a complex, highly organized living organ undergoing continuous remodeling throughout life. It contains a large amount of organic material (40%). Some 90–96% of this organic material is collagen, predominantly type 1 along with small amounts of types V and XII; the remainder (4–10%) is bone specific proteoglycans and noncollagenous proteins such as osteocalcin, osteonectin, bone sialoprotein, and bone phosphoproteins. Most of the inorganic matrix or mineral phase of the bone is the crystal of an apatite of calcium and phosphate. This gives bone stiffness and strength. Within this crystal structure other mineral ions such as citrate, carbonate, fluoride, and hydroxyl ions also exist. Bone plays a very important role as a primary reservoir of calcium in the body and exchanges this mineral readily with the extracellular fluid environment. Also, the hematopoietic marrow, which supplies nutrient-carrying red blood cells and infectionfighting white blood cells, is located in the trabecular bone. Another function of bone is its 301

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mechanical role in supporting the body’s tissues and in providing sites for attachment for the muscles that affect body movement and locomotion. Bones can be classified according to their shape in three groups: short, flat, and long bones. Short bones such as tarsals, carpals, and vertebral bodies have thin cortices and are trapezoidal, cuboidal, or irregular in shape. Flat bones have one dimension. The larger flat bones form the cranial vault, the wing of the ileum, and the scapula, whereas the smaller ones form the lamina of a vertebra. Long bones or tubular bones such as femur, tibia, humerus, metacarpals, metatarsals, and phalanges are usually made of spongy and compact bone. The spongy bone consists of three-dimensional branching of trabeculae interdispersed by bone marrow. The spongy bone changes gradually into compact bone toward the middle of the bone. Cortical (compact) and cancellous (trabecular) bones are the two important bone tissues having the same matrix composition and structure. Cortical bone has greater density or less porosity than cancellous bone. Cortical bone forms approximately 80% of the mature skeleton and surrounds the marrow and the cancellous bone plates. In long bones, cortical bones form the diaphysis [1]. The metaphyses contain most of the cancellous bone. The cortical bone decreases in thickness from the middle part of the diaphysis to the metaphysis where the cancellous bones are arranged to support the subchondral bone. In long bones, cortical bone provides the maximal resistance to torsion and bending. However, cancellous bone allows great deformation under the same load. Cortical bone also protects the articular cartilage and subchondrial bone from damage by absorbing the impact loads applied across synovial joints. Cancellous bone has approximately 20 times more surface area per unit of volume than does cortical bone and has a higher rate of metabolic activity and remodeling. It can respond more rapidly to changes in mechanical loads than does cortical bone. The external and internal surface of the bone is called periosteum and endosteum, respectively, and both have osteogenic properties. Periosteum is primarily affected from diseases, deformities, injuries of the skeleton, and most orthopedic treatments. It contributes an important part of the blood supply to the bone. Periostal cells can resorb and form bone in response to local or systemic stimuli. During bone growth, they secrete the organic matrix that enlarges the diameter of the bone. Some tendons and ligaments are inserted primarily into the outer layer of the periosteum. Bone marrow is found in the internal part of the bone and serves as a source of bone cells. The blood vessels in the marrow form a critical part of the circulatory system in bone. Bone cells originate from a mesenchymal stem cell line [e.g., undifferentiated cells (preosteoblasts), osteoblasts, bone lining cells, and osteocytes] and a hematopoietic stem cell line (e.g., marrow monocytes, preosteoclasts, and osteoclasts). Osteoblasts are differentiated cells, usually derived from osteoprogenitor cells. They are mobile and change shape and size with different rates of matrix production. They appear with the vascular tissue. The main function of the osteoblasts is the synthesis and secretion of the organic matrix of bone. They also control electrolyte fluxes between the extracellular fluid and the osseous fluid and may influence the mineralization of bone matrix through the synthesis of organic matrix components of bone and the production of matrix vesicles. Moreover, some systemic hormones such as parathyroid hormone and local cytokines may stimulate osteoblasts to release mediators that activate osteoclasts [2]. Bonelining cells lie against the bone matrix and when exposed to parathyroid hormone, they contract and secrete enzymes that remove the thin layer of osteoid. This event permits osteoclasts to attach and resorb the surface of bone. Osteocytes are the most abundantly found bone cells in the mature human skeleton. They surround themselves with an organic

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matrix and cover the matrix together with the periostal and endosteal cells. Osteoclasts have a very efficient method for destroying bone matrix. They are rarely found in normal bone. Osteoclast precursor cells are found in the marrow or the circulating blood. Stem cells are activated by specific hormones and growth factors to become osteoclast precursors. When osteoclasts are active, they require a great amount of energy to resorb bone. Bone resorption and bone formation are the two main events of the remodeling process, with two well-differentiated cell types, osteoclasts, osteoblasts, and their precursors. The first step is activation that involves resting osteoblasts or lining cells as well as osteoclast precursors. In the resorption phase, the osteoclasts remove bone to produce Howship’s lacunae in trabecular bone or Haversian canals in cortical bone. After osteoclastic resorption is complete, mononuclear cells are seen in the reversal phase and may form the cement line. Thereafter, osteoblast precursors migrate to the bone surface and lay down new matrix in the formation phase. Bone cell function is regulated at both the systemic and the local level. Factors that control local formation and resorption of bone such as cytokines, prostaglandins, and mechanical loading have the potential for the treatment of acute fractures, delayed unions, nonunions, and skeletal deformity as well as to enhance the stabilization of joint implants in the skeleton. II CONVENTIONAL THERAPY OF BONE TISSUE ENGINEERING Organ or tissue failure or loss is an expensive problem in health care. When the mechanical function of bone is lost by injury, infection, disease, or inadequate fracture management, it can only be regained by restoring skeletal continuity at the location of interest. In these situations, the patient suffers from the loss of mechanical support of the body’s tissues and loss of muscle attachment sites. Such skeletal defects, therefore, require therapeutic intervention to restore the bone to its proper form. Currently, major approaches to tissue or organ loss are surgical reconstruction, transplantation, and artificial prosthesis. Although these therapies have saved and improved countless lives, they remain imperfect solutions. Surgical reconstruction and artificial prosthesis can result in infection, chronic irritation, lack of mechanisms of biological repair, and long-term problems such as development of malignant tumor or calcification of graft. Transplantation is currently limited by increasing donor shortages. Also immunosuppression is required with its serious side effects. Existing autogenous graft materials may play an important role in healing nonunion defects. However, they generally have poor morphologies and interconnectivities that are dissimilar to cancellous bone. They also have limited resource that leads to donor site morbidity. This may impair their population with cells and vascularization. Allograft bone can also be a successful procedure. Allograft bone has a porous structure and contains some growth factors such as insulin-like growth factor, transforming growth factor-, plateletderived growth factor, fibroblast growth factor, and bone morphogenetic proteins (BMPs) embedded in the bone matrix. These growth factors make allograft bone potentially osteoinductive [3]. Allograft bone is fully resorbable and biocompatible and does not have reactive degradation byproducts. However its high cost, risk of bacterial or viral infection, and donor-to-donor variation in quality are major disadvantages of allograft bone. Various synthetic alternatives such as metals, ceramics, and polymers have been tried for many years. However, they have had limited success due to lack of living cells and slow vascularization.

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A Ceramics Hydroxyapatite and tricalcium phosphate (TCP) are the most widely used ceramic materials. There are some drawbacks of ceramics usage. Ceramics are brittle and have poor tensile strength. Their use in clinical situations requiring significant torsional, impact, or shear stress is limited. The rate of bioresorption of ceramics depends on their chemical composition. Hydroxyapatite is a slowly resorbing calcium phosphate ceramic. Due to the insoluble and inert structure of crystalline hydroxyapatite, remodeling is extremely slow. Large amounts of hydroxyapatite may remain in the body for 10 years [4]. It does not lead to immunologic reactions. It was shown earlier that hydroxyapatite implants were invaded by fibrovascular tissue and then converted to mature lamellar bone [5]. This newly formed bone was nearly identical to that seen in autogenous grafts. Holmes et al. [6] also showed resorption of the surface of the hydroxyapatite implant by osteoclast-like cells. Because the structure of hydroxyapatite is inert and insoluble, remodeling was extremely slow. They did not observe any immunological or giant cell reaction surrounding implants. Because hydroxyapatite implants are brittle and remodel slowly, they have not been recommended for diaphyseal defects [3]. They would leave diaphyseal sites susceptible to fatigue fracture through the residual implant or to fracture at the junction between the implant and the host bone. In order to increase the rate of resorption, several modifications have been tested. One of them is a composite of hydroxyapatite and calcium carbonate [3]. In this composite, the material is almost entirely calcium carbonate coated with a thin layer of hydroxyapatite. Another modification is a composite of hydroxyapatite (30%) and the more soluble TCP (70%) [7,8]. Implantation of this material gave the new bone formation a more stable and long lasting framework. Tricalcium phosphate is the rapidly resorbable calcium phosphate ceramic and is reabsorbed 10 to 20 times faster than hydroxyapatite. Because the porosity of bulk TCP implants is too small, bone ingrowth within the matrix material becomes difficult [9]. Due to increased porosity of the matrix and the bioavailable surface, granules of TCP may be more effective than bulk TCP. Frost [10] showed that the presence of TCP may stimulate local osteoclasts, which in turn stimulate osteoblastic activity with new bone formation. Injectable calcium phosphate cement containing -TCP, dibasic dicalcium phosphate, and tetracalcium phosphate monoxide was used by Jupiter et al. [11]. They showed that these cements are useful for treatment of distal radius fractures. Calcium sulfate hemihydrate (plaster of Paris) was used by Turner et al. [12] as a synthetic graft material. They reported that it was absorbed by osteoclasts in several animal studies. In contrast to other ceramics, host bone formed in concentric layers around the resorbing plaster of Paris. B Demineralized Bone Matrix Demineralized bone matrix (DBM) is a potent inducer of bone formation. It contains a bone-inducing substance that causes the responding mesenchymal cells to undergo differentiation into chondroblasts, osteoblasts, and hematocytoblasts with the resultant formation of endochondral bone. It also retains growth factors and noncollagenous bone proteins such as osteocalcin and osteopontin. Demineralized bone matrix can be used alone or in combination with autogenous bone grafts or ceramics [13]. Urist [14] implanted the demineralized bone into muscle pouches in rodents and discovered that it stimulates the formation of ectopic bone. Glowacki et al. [15] used DBM in

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craniofacial reconstruction with success. It was also used efficiently in osseous defects such as bone defects in children, nonunited fractures, comminuted fractures, and arthrodeses [16]. Polydioxanone has been combined in semirigid sheet form with demineralized bone matrix in rat calvarial defects. It functioned as a restraining agent for the particulate DBM and provided shape control [17]. C Hydrogels Mooney and Mikos [18] have studied the potential of injectable, biodegradable hydrogels—gelatin-like, water-filled polymers—for treating dental defects, such as poor bonding between teeth and the underlying bone, through guided bone regeneration. The hydrogels can be made to incorporate molecules that both modulate cellular function and induce bone formation. They provide a scaffold on which new bone can grow and minimize the formation of scar tissue within the regenerated region. D Polyurethanes Polyurethane was used in membrane form to construct a tube around a rabbit radius defect [19]. Ten out of ten defects healed, compared to one out of ten untreated controls. This is an example of the guided tissue regeneration principle, in which the tube of polymer is purported to prevent the ingress and ingrowth of tissues different from the injured tissue during healing. Saad et al. [20] showed the chondrocyte and osteoblast compatibility on biodegradable polyesterurethane foam. During the degradation of this polymer, small crystalline particles of short chain poly-(R)-3-hydroxybutyric acid (PHB-P) and lysine methyl ester were released, and osteoblasts showed only limited PHB-P phagocytosis or no signs of any cellular damage at low concentrations. At high concentrations of PHB-P, the cell viability of macrophages and to a lesser extent of osteoblasts was affected. E Poly(Propylene Fumarate) Poly(propylene fumarate) (PPF) has been incorporated into composite materials to replace trabecular bone local in high concentrations [21]. The fumaric acid double bond of polymer chain has been used to secondarily crosslink the PPF during formulation of the composite material [22]. This material has demonstrated adequate mechanical properties in vitro to function as a temporary trabecular bone replacement [23]. F Polylactic Acid, Polyglycolic Acid, and Copolymers Poly(-hydroxy esters) are promising substrates for osteoblast transplantation and have been extensively used in tissue engineering studies because of their biocompatibility, their safety as suture material, and their potential for use as a time-release delivery system for peptide growth factors. Three types currently under investigation for this type of application are poly(L-lactic acid) (PLLA), poly(glycolic acid) (PGA), and poly(lactic-co-glycolic acid) (PLGA). Their degradation occurs by passive hydrolysis and the degradation rate can be influenced by water uptake, pH, bond type, crystallinity, and steric hindrance. They undergo nonenzymatic hydrolysis of the polyester bond resulting in either lactic or glycolic acid [24]. Release of these degradation products can cause a decrease in the local pH, and at high concentrations results in tissue damage. Pore sizes, porosity, and degradation rates can be controlled and varied by adjusting processing parameters.

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The early studies with those polymers were not very satisfactory. They are well tolerated in the early phase after implantation, but an inflammatory response is noted in tissues surrounding implants [3]. Bostman et al. [25] reported an adverse foreign body reaction in nearly all clinical studies with PLGA fracture fixation screws. Among the poly(-hydroxy esters), PLLA has the highest mechanical strength and was investigated for orthopedic applications in the form of pins, screws, and plates [26]. Although PLLA proved to be biocompatible, there is great concern regarding the toxicity of degradation products, which include lactic acid and undegraded PLLA particles. The amount of lactic acid released during the course of PLLA degradation is very small, but increases rapidly as PLLA is broken down to low molecular weight oligomers. This event leads to a decrease in the local pH, which in turn may accelerate the PLLA degradation rate and induce an inflammatory reaction. This problem may be overcome by using polydispersed PLLA, though doing so may compromise the mechanical properties [27]. The use of polydisperse PLLA results in a distribution of lactic acid production more evenly over the period of degradation. The formation of undegraded PLLA microparticles due to the difference in degradation rates between the crystalline and amorphous regions of PLLA may also trigger foreign body inflammatory reactions such as the bone resorption observed with PLLA implants. III BONE GROWTH FACTORS A Types of Bone Growth Factors Growth factors when combined with implants would yield a much faster and better tissue repair. Some growth factors play a critical role in bone healing and regeneration. These tissue-specific polypeptides act as local regulators of cellular activity. They bind to cell surface transmembrane receptors on the target cell and intracellular domain is stimulated. This event usually causes an activation of specific protein kinases leading to activation of transcription of a gene into mRNA and translation of proteins for use intracellularly or extracellularly [28]. These growth-promoting factors include BMPs, insulin-like growth factors (IGF), transforming growth factor- (TGF-), fibroblast growth factors (FGF), and platelet-derived growth factors (PDGF). 1 Bone Morphogenetic Proteins Bone Morphogenetic Proteins appear to be the most promising low molecular weight peptides. They initiate bone formation by recruiting and stimulating local progenitor cells of osteoblast lineage and by enhancing bone collagen synthesis [29]. They belong to an expanding TGF- superfamily. They play an important role in the formation and maturation of skeletal tissues. They can be extracted from bone, dentine, and osteosarcomata or metastases of prostate cancer. More than 24 have been identified. The most studied and osteoinductive are BMP-2, BMP-4, and BMP-7. They stimulate alkaline phosphatase activity and collagen synthesis by osteoblastic cells in vitro. Bone morphogenetic protein-7 (OP-1) promotes chondroblastic and osteoblastic differentiation and formation of mineralized matrix by osteoblasts. 2 Transforming Growth Factor- Transforming growth factor-, a chemotactic mediator for fibroblasts and macrophages, blocks plasminogen inhibitor, increases angiogenesis, and stimulates collagen synthesis,

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mesenchymal cell growth, differentiation, and bone repair. It inhibits procollagenase and epithelial expansion and also increases the osteoinductive activities of the BMPs. Moreover, TGF- regulates the proliferation of chondrocytes and synthesis of a cartilage matrix. There are three isoforms of TGF-. Type 1 is stored in platelet alpha granules and plays a crucial role in fracture repair. 3 Platelet-Derived Growth Factor Platelet-derived growth factor is synthesized by platelets, macrophages, monocytes, and endothelial cells. It increases DNA synthesis, cell replication, production of collagen, and noncollagenous proteins. Also, plays a role in fracture repair and osteoclast activation. 4 Insulin-Like Growth Factor Insulin-like growth factor promotes cellular proliferation and matrix synthesis by osteoblasts and chondroblasts. The most abundant isomer in human bone cells is IGF-II. 5 Fibroblast Growth Factors Fibroblast growth factors act on stem and differentiated cells, causing changes in migration, proliferation, differentiation, morphology, and function. They are mitogenic factors that are produced in an early phase of fracture repair and localized in the bone matrix. Acidic and basic fibroblast growth factors induce and inhibit bone repair, depending on the applied dose. B Bone Growth Factor Delivery Bone growth factors play an important role in bone tissue regeneration and repair. However, most of them have short half-lives after intravenous injection. Therefore, their biological activity lasts only a few minutes in the circulation and repeated injection is required to obtain sustained blood levels. These molecules are also large and extremely potent; therefore, they penetrate the tissue barriers very slowly and their systemic administration can lead to toxicity. In order to solve these problems, new delivery methods are required for growth factors. Polymers can be designed so that the scaffold material provides both structural and controlled release function. Bone morphogenetic protein was delivered to a segmental rabbit tibial defect via a polymer containing PLLA, PEG, and PLGA [30]. The researchers reported a complete restoration of the defect in 12 weeks. Yasko et al. [31] used rat DBM as a carrier and delivered BMP-2 to a rat femoral defect. They reported complete healing of the defects in this study. Bone morphogenetic protein was delivered to treat nonunions in a PLLA-PLGA delivery system [32]. All patients healed their nonunions within 5 months after the treatment. Bone morphogenetic proteins were also carried with PLGA-TCP composite [33], collagen [34], plaster of Paris [35], and hydroxyapatite impregnated with collagen [36]. These systems all resulted in the formation of new bone in heterotopic sites, and more bone with increasing BMP doses was reported. The addition of bioactive molecule induces bone growth to occur faster and in greater mass. Liebergall et al. [37] incubated cells in TGF-1 and delivered on ceramic to the defect site. They observed increase in bone growth. Summer et al. [38] used titanium rods plasma flame–sprayed with hyaluronic acid (HA) and TCP for canine bilateral implantation. They treated the implants with TGF-1 and observed increase in bone growth. Bovine collagen was used as a TGF- vehicle because it is highly purified, hypoallergenic, and can

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form a viscous suspension that can easily be applied to the cut edges of an incision [39]. Slow release was reported and prolonged local exposure of TGF-. Strates et al. [40] observed large amount of bone formation in a nonunion study in dogs by using TGF- adsorbed on resorbable microcrystals of hydroxyapatite. An application of human recombinant TGF-1 in a 3% methylcellulose gel to skull defects created in rabbits induced a dose-dependent increase in intramembranous bone formation [41]. Delivery of TGF-1 by PLGA and DBM was also used [42]. Most of the TGF1 released from the delivery system retained its bioactivity, and the devices were sufficiently porous to allow bone ingrowth. IV BONE TISSUE ENGINEERING USING BIODEGRADABLE SCAFFOLDS Bone tissue engineering can be addressed to solve problems such as high cost, risk of bacterial or viral infection, donor shortage, donor-to-donor variation in quality, and slow vascularization, etc. For loss of bone or its function, ready-made pliable tissue constructs are very promising. For a successful application of such tissues, the development of biomaterials with suitable mechanical characteristics and degradation behavior is very important. To minimize a loss of cells, it might be advantageous to preform the biomaterial matrix to a needed shape. Osteoconductive matrices are fashioned from biodegradable materials of natural origin like collagen, gelatin, hyaluronic acid, etc., biodegradable polymers like poly(3-hydroxybutyric acid-co-3-hydroxyvaleric acid) (PHBV), or most commonly from synthetic polymers such as polylactic acid, polyglycolic acid, poly(lactic-co-glycolic acid), etc. The matrices used as scaffolds should satisfy certain requirements. They should be designed to allow diffusion of nutrients to the transplanted cells and guide cell organization, attachment, and migration. They also should be biodegradable and bioresorbable. In addition, scaffold materials should reduce immune response to allogenic cells. Pore size and porosity are other important parameters. They have to allow optimal cell invasion, fibrovascular tissue, and new blood vessels. The mechanical properties of the material should match those of the tissue. Mechanical properties of scaffolds during degradation and bone regeneration must be characterized. There is increasing evidence that changes in scaffold surface chemistry and topography alter cellular activity [43]. Therefore, surfaces may need to be characterized or even altered to facilitate bone tissue regeneration. A Hyaluronic Acid Hyaluronic acid is found in normal joint fluid, early fracture callus, and hyaline cartilage. It is generally used in the eye to replace the vitreous humor and can also be used for intraarticular injection in the knee in early osteoarthritis. The degradation rate of hyaluronic acid can be modulated through covalent crosslinking. Solchaga et al. [44] used two biomaterials (HYAFF 11 and ACP) based on hyaluronic acid modified by esterification of the carboxyl group of the glucuronic acid. Those materials were tested as osteogenic or chondrogenic delivery vehicles for rabbit mesenchymal progenitor cells and compared with a well-characterized porous calcium phosphate ceramic delivery vehicle. They found that the hyaluronic acid–based delivery vehicles are superior to porous calcium phosphate ceramic in terms of the number of cells loaded per unit volume of implant, and HYAFF 11 sponges are superior to the ceramics with regard to the amount of bone and cartilage formed. Additionally, hyaluronic

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acid–based delivery vehicles have the advantage of degradation/resorption characteristics that allow complete replacement of the implant with newly formed tissue. B Collagen Collagen is the most abundant single protein in most vertebrates. It is the matrix on which the mineral constituents precipitate. It forms the major portion of tendons and it is an important constituent of skin. Collagen is characterized by its typical triple-helical domain. About 90% of human collagen is type I, found mainly in skin and bone, whereas collagen type II is the major constituent of cartilage. Collagen contains information such as particular amino acid sequences that may facilitate cell attachment or maintenance of differentiated function. However, a high rate of degradation is the major drawback of this material. In order to prevent this rapid degradation, collagen can be incorporated with calcium phosphate to create a bone-mimicking material. A nano-HAp/collagen (nHAC) composite that mimics natural bone both in composition and microstructure to some extent was employed as a matrix for the tissue engineering of bones [45]. The porous nHAC scaffold provided a microenvironment resembling that seen in vivo, and cells within the composite eventually acquired a tridimensional polygonal shape. In addition, new bone matrix was synthesized at the interface of bone fragments and the composite. Hsu et al. [46] used microspheres of hydroxyapatite/reconstituted collagen as support for osteoblast cell growth and found that these microspheres could be used as the filling material for bone defects. Cornell et al. [47] compared Collagraft plus autogenous marrow versus cancellous iliac bone grafts in acute long bone fractures. Collagraft is a commercially synthesized composite of suspended, deantigenated bovine fibrillar collagen and porous calcium phosphate ceramic (65% hydroxyapatite, 35% TCP). It is nonosteoconductive and the addition of autogenous bone marrow provides osteoprogenitor stem cells. At the end of the study, no significant or radiographic differences were found. The use of Collagraft shortened operative time and avoided the operative morbidity of harvesting iliac crest bone grafts. C Polylactic Acid, Polyglycolic Acid, and Their Copolymers Osteoblasts cultured on PLGA films were found to attach, proliferate, migrate, and secrete collagen and alkaline phosphatase (ALPase) at rates comparable to tissue culture polystyrene [48]. Long-term three-dimensional in vitro studies of porous PLGA foams also demonstrated potential for bone regeneration [43]. In an in vivo study, a rat nonunion defect was treated with periosteal cell/polymer construct, chondrocyte/polymer construct, polymer without cells, or nothing [47]. The polymer scaffold was a nonwoven mesh of PGA fibers. This study strongly suggests that polymer constructs delivering bone cells to the defect site have osteoinductive potential. Ishaug et al. [50] demonstrated that osteoblasts on polymer films migrate from isolated osteoblast cultures and bone chips as a monolayer of cells and continue to proliferate to form multilayers. They also reported [43] the formation of ectopic bone in rat mesentery by using PLGA foams. D Poly(3-Hydroxybutyric Acid-co-3-Hydroxyvaleric Acid) Poly(3-hydroxybutyric acid-co-3-hydroxyvaleric acid) is an optically active polyester consisting of D--hydroxybutyric acid (PHB) as a repeating unit, which is produced by mi-

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croorganisms and is a normal constituent of human blood. It is produced by various strains of microorganisms such as Bacillus megaterium, a glucose-utilizing mutant of Alcaligenes eutrophus, etc., soil bacteria, estuarine microflora, blue-green algae, and microbiologically treated sewage. In microorganisms, PHB serves as an intracellular energy and carbon storage product in much the same way as glycogen in mammalian tissue. It is also thought to have a minor role in cellular function such as sporulation, encystment, and gene expression. The polymer accumulates in discrete membrane-bound granules in the bacterial cells from which it can be extracted directly with organic solvents such as chloroform or by membrane rupturing techniques. It is safe, nontoxic, biocompatible, biodegradable, easily and reproducibly processable. The related copolymers of PHB with -hydroxyvaleric acid (PHBV) emerge as a new generation of PHB-based materials with more adjustable properties depending upon copolymer composition. Degradation rate of PHBV can easily be adjusted by changing the copolymer composition. Matrices of PHB and PHBV lose mass very slowly when compared to bulk-degrading poly(lactide–glycolide) systems. Poly(hydroxybutyricco-hydroxyvaleric acid) copolymers are semicrystalline and have high degrees of crystallinity (60–80%). Crystallinity influences polymer properties such as rate and mechanism of degradation, drug compatibility, and drug release. They are also biocompatible [51,52]. Further they are known to exhibit piezoelectricity [53]. Since electrical stimulation is thought to promote bone healing and repair, polymers have been suggested for use as bone pins and plates. Rivard et al. [54] demonstrated that PHBV (9%) sustained a fibroblast cell proliferation rate similar to that observed in collagen sponges for up to at least 35 days. On the other hand, the PHBV materials kept their integrity during the culture period, while the collagen foams contracted greatly. Moreover, the total protein production after 4 weeks in culture was found to be twice as high in the PHBV foam than in the collagen foam. Porous PHBV materials appear to be adequate polymeric substrates for cell cultures. V BONE TISSUE ENGINEERING WITH OSTEOBLAST-SEEDED PHBV8 We also have chosen to work with PHBV8 matrix material. In this study, foams were prepared from PHBV8 solutions containing 4, 6, and 8% w/w polymer. In order to increase the void volume, polymer solutions were loaded with sucrose crystals that were leached out after the formation of the membrane. Their surfaces were treated with oxygen rf-plasma to modify their surface chemistry and hydrophilicity. Scanning electron microscopic examination of the PHBV matrices was carried out. It was observed from the SEM of the untreated PHBV that as the polymer concentration increased from 4 to 8%, a more ordered and compact surface was obtained. In the case of oxygen rf-plasma–treated PHBV-sieved sucrose foams, a more porous structure than that of an untreated one was obtained (Figs. 1–3). Changes in pH of the aging media with time and weight loss of unloaded and PHBVsieved sucrose of 4, 6, and 8% foams were determined. After 180 days there was nearly no change in the weight of the unloaded PHBV polymers or the pH of the medium. However, a pH change was observed in the medium where PHBV-sieved sucrose foams were placed. The pH decrease is an indication of the degradation extent. Osteoblast cells were isolated from rat bone marrow. Because culture conditions have been shown to affect osteoblast activity [55,56], some medium supplements such as L-ascorbic acid, -glycerophosphate, and dexamethasone were used for osteoblast growth.

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Figure 1 SEM of unloaded PHBV8 (6%).

Osteoblasts obtained from fetal calvaria secreted mineralized matrix when L-ascorbic acid and -glycerophosphate were added to the medium. L-Ascorbic acid is an enzyme cofactor and antioxidant that stimulates the transcription, translation, and posttranslational processing of collagen in connective tissue cells. In culture of bone-derived cells, ascorbate stimulates osteoblastic differentiation, synthesis, and deposition of collagen and mineralization. Ascorbate influences the differentiation of preosteoblasts and is required for the synthesis of osteoid by mature osteoblasts. Ascorbate and -glycerophosphate are the inducers of osteogenic differentiation. In the presence of -glycerophosphate, bone matrix deposition is induced and increased with time in culture. Dexamethasone, a glucocorticoid, increased the number of collagenous mineralized nodules found in calvaria culture and extended the days that nodules formed in culture. Stromal cells obtained from bone marrow also secreted matrix with collagenous, mineralized nodules, but only when the medium was supplemented with both -glycerophosphate and dexamethasone. Differentiated cultured bone cells provide an essential tool for investigating biodegradable scaffolds for bone tissue regeneration.

Figure 2 SEM of PHBV8 (6%) with sieved sucrose.

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Figure 3 SEM of oxygen plasma–treated PHBV8 (6%) with sieved sucrose. In addition, enhancement of osteoblast differentiation on scaffolds prior to transplantation may improve effectiveness of this type of bone regeneration therapy. After isolation of osteoblasts, they were observed under a light microscope and microphotographs were obtained. Then they were seeded onto PHBV8 matrices. In order to determine the cell density on the foams, a Cell Titer 96™ Nonradioactivity Cell Proliferation Assay (MTS) test was conducted; MTS was also used to construct a calibration curve with different concentrations of osteoblasts. Results showed that osteoblasts proliferated on PHBV8. The exact amount of the cells after 7 days of incubation on the PHBV films was determined from the calibration curve. It was found that cells grew better on large pore size foams (made by 300–500 m sucrose crystals) than on the small pore size foams (75–300 m sucrose crystals). A cell concentration of 5.105 cells/mL seemed to be the ideal loading. Production of alkaline phosphatase (ALP), an enzyme specific for bone, was measured spectrophotometrically. According to the results, PHBV8 seems to be a promising polymeric matrix for bone tissue engineering. REFERENCES 1. Buckwalter J. A., Glimcher M. J., Cooper R. R., Recker R. 1995. Bone biology: structure, blood supply, cells, matrix, and mineralization. J. Bone Joint Surg. 77A:1256–1275. 2. Buckwalter J. A., Glimcher M. J., Cooper R. R., Recker R. 1995. Bone biology: formation, form, modeling, remodeling, and regulation of cell functions. J. Bone Joint Surg. 77A:1276– 1287. 3. Fleming J. E., Cornell C. N., Muschler G. F. 2000. Bone cells and matrices in orthopedic tissue engineering. Orthop. Clin. N. Am. 31:357–374. 4. Gazdag A. R., Lane J. M., Glaser D. 1995. Alternatives to autogenous bone graft: efficacy and indications. J. Am. Acad. Orthop. Surg. 3:1–8. 5. Chapman M. W., Bucholz R., Cornell C. N. 1997. Treatment of acute fractures with a collagen–calcium phosphate graft material: a randomized clinical trial. J. Bone Joint Surg. Am. 79:495–502. 6. Holmes R., Bucholz R., Mooney V. 1986. Porous hydroxyapatite as a bone graft substitute in metaphyseal defects. J. Bone Joint Surg. Am. 68:904–911. 7. Zerwekh J. E., Kourosh S., Schienbergt R. 1992. Fibrillar collagen–biphasic calcium phosphate composite as a bone graft substitute for spinal fusion. J. Orthop. Res. 10:562–572.

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8. Buckholz R. W. 1987. Clinical experience with bone graft substitutes. J. Orthop. Trauma 3:260–262. 9. Goldberg V. M., Stevenson S., Shaffer J. W. 1991. Biology of autografts and allografts. In: Bone and Cartilage Allografts. American Academy of Orthopedic Surgery: Park Ridge, IL, pp. 3–13. 10. Frost H. 1991. A new direction for osteoporosis research: a review and proposal. Bone 12:429–437. 11. Jupiter J. B., Winters S., Sigman S. 1997, Repair of five distal radius fractures with an investigational bone cement: a preliminary report. J. Orthop. Trauma 11:110–116. 12. Turner T. M., Urban R. M., Gitelis S. 1991. Efficacy of calcium sulphate, a synthetic bone graft material, in healing a large canine medullary defect. Trans. ORS 24:522. 13. Khan S. N., Tomin E., Lane J. M. 2000. Clinical applications of bone graft substitutes. Orthop. Clin. N. Am. 31:389–398. 14. Urist M. R. 1965. Bone: formation by autoinduction. Science 150:893–899. 15. Glowacki J., Mulliken J. B. 1985. Demineralized bone implants. Clin. Plast. Surg. 12:233. 16. Tiedeman J. J., Garvin K. L., Kile T. A. 1995. The role of composite, demineralized bone marrow in the treatment of osseous defects. Orthopedics 18:1153–1158. 17. Nichter R. S., Yazdi M., Kosari K., Sridjaja R., Ebramzadeh E., Nimni M. E. 1992. Demineralized bone matrix polydioxanone composite as a substitute for bone graft: a comparative study in rats. J. Craniofac. Surg. 3:63–69. 18. Mooney D. J., Mikos A. G. 1999. Growing new organs. Scientific American 280(4):60–65. 19. Nielsen F., Karring T., Gogolewski S. 1992. Biodegradable guide for bone regeneration. Polyurethane membranes tested in rabbit radius defects. Acta Orthop. Scand. 63:66–69. 20. Saad B., Neunenschwander P., Uhlschmid G. K., Suter U. W. 1999. New versatile, elastomeric, degradable polymeric materials for medicine. Int. J. Biol. Macromol. 25:293–301. 21. Gerhart T. N., Roux R. D., Horowitz G., Miller R. L., Hanff P., Hayes W. C. 1988. Antibiotic release from an experimental biodegradable bone cement. J. Orthop. Res. 6:585–592. 22. Hasirci V., Lewandrowski K.-U., Bondre S. P., Gresser J. D., Trantolo D. J., Wise D. L. 2000. High strength bioresorbable bone plates: preparation, mechanical properties, and in vitro analysis. Biomed. Mater. Eng. 10:19–29. 23. Yaszemski M. J., Payne R. G., Hayes W. C., Langer R., Mikos A. G. 1996. The in vitro degradation of a poly(propylene fumarate)–based composite material. Biomaterials 17:2127–2130. 24. Behravesh E., Yasko A. W., Engel P. S., Mikos A. G. 1999. Synthetic biodegradable polymers for orthopedic applications. Clin. Orthop. Rel. Res. 367S:S118–S125. 25. Bostman O., Hirvensalo E., Markmen J. 1990. Foreign body reactions to fracture fixation implants of biodegradable synthetic polymers. J. Bone Joint Surg. Br. 2:592–596. 26. Rader C. P., Keller H. W., Rehm K. E. 1992. Surgical treatment of dislocated 3- and 4-segment fractures of the proximal humerus. Unfallchirurg 95:613–617. 27. von Recum H. A., Cleek R. L., Eskin S. G., Mikos A. G. 1995. Degradation of polydispersed poly(L-lactic acid) to modulate lactic acid release. Biomaterials 16:441–447. 28. Solheim E. 1998. Growth factors in bone. Int. Orthop. 22:410–416. 29. Khan S. N., Bostrom M. P. G., Lane J. M. 2000. Bone growth factors. Orthop. Clin. of N. Am. 31:375–387. 30. Miyamoto S., Takaoka K. 1993. Bone induction and bone repair by composites of bone morphogenetic protein and biodegradable synthetic polymers. Ann. Chir. Gynaecol. 82:69–75. 31. Yasko A. W., Lane J. M., Fellinger E. J., Rosen V., Wozney J. M., Wang E. A. 1992. The healing of segmental bone defects induced by recombinant human bone morphogenetic protein (rhBMP-2): a radiographic, histological, and biomechanical study in rats. J. Bone Joint Surg. 74A:659–670. 32. Johnson E. E., Urist M. R., Finerman G. 1990. Distal metaphyseal tibial non-union: deformity and bone loss treated by open reduction, internal fixation, and human bone morphogenetic protein (hBMP). Clin. Orthop. Rel. Res. 250:234–240.

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33. Desilets C. P., Marden L. J., Patterson A. L., Hollinger J. O. 1990. Development of synthetic bone-repair materials for craniofacial reconstruction. J. Craniofac. Surg. 1:150–153. 34. Gao T. J., Lindholm T. S., Marttinen A., Puolakka T. 1993. Bone inductive potential and dosedependent response of bovine bone morphogenetic protein combined with type IV collagen carrier. Ann. Chir. Gynaecol. 82:77–84. 35. Yamazaki Y., Oida S., Akimoto Y., Shioda S. 1988. Response of the mouse femoral muscle to an implant of a composite of bone morphogenetic protein and plaster of Paris. Clin. Orthop. Rel. Res. 234:240–249. 36. Takaoka K., Nakahara H., Yoshikawa H., Masuhara K., Tsuda T., Ono K. 1988. Ectopic bone induction on and in porous hydroxyapatite combined with collagen and bone morphogenetic protein. Clin. Orthop. Rel. Res. 234:250–254. 37. Liebergall M., Young R. G., Ozawa N., Reese J. H., Davy D. T., Goldberg V. M., Caplan A. I. 1994. The effects of cellular manipulation and TGF- in a composite bone graft. In: Bone Formation and Repair, Brighton C. T., Friedlaender G. E., Lane J. M., Eds. American Academy of Orthopedic Surgeons: Rosemont, IL. 38. Sumner D. R., Turner T. M., Purchio A. F., Gombotz W. R., Urban R. M., Galante J. O. 1995. Enhancement of bone ingrowth by TGF-. J. Bone Joint Surg. 77A:1135–1147. 39. Beck L. S., DeGuzman L., Lee W. P., Xu Y., Siegel M. W., Amento E. P. 1993. One systemic administration of TGF-1 reverses age- or glucocorticoid-impaired wound healing. J. Clin. Invest. 92:2841–2849. 40. Strates B. S., Kilaghbran V., Nimni M. E., McGuire M. H. 1992. Enhanced periostal osteogenesis induced by TGF-1 adsorbed on microcrystals of hydroxyapatite. Trans. Orthop. Res. Soc. 591. 41. Beck L. S., DeGuzman L., Lee W. P. 1991. TGF-1 induces bone closure of skull defects. J. Bone Min. Res. 6:1257–1265. 42. Gombotz W. R., Pankey S. C., Ranchalis L. S., Puolakkainen P. 1993. Controlled release of TGF-1 from a biodegradable matrix for bone regeneration. J. Biomater. Sci. Polym. Ed. 5(1–2):49–63. 43. Ishaug S. L., Crane G. M., Miller M. J., Yasko A. W., Yaszemski M. J., Mikos A. G. 1997. Bone formation by three-dimensional stromal osteoblast culture in biodegradable polymer scaffolds. J. Biomed. Mater. Res. 36:17–28. 44. Solchaga L. A., Dennis J. E., Goldberg V. M., Caplan A. I. 1999. Hyaluronic acid–based polymers as a cell carriers for tissue engineered repair of bone and cartilage, J. Orthop. Res. 17:205– 213. 45. Du C., Cui F. Z., Zhu X. D., de Groot K. 1999. Three dimensional nano-HAp/collagen matrix loading with osteogenic cells in organ culture. J. Biomed. Mater. Res. 44:407–415. 46. Hsu F. Y., Chueh S. C., Wang Y. J. 1999. Microspheres of hydroxyapatite/reconstituted collagen as supports for osteoblast cell growth. Biomaterials 20:1931–1936. 47. Cornell C. N., Lane J. M., Chapman M. 1991. Multicenter trial of Collograft as bone graft substitute. J. Orthop. Trauma 5:1–8. 48. Ishaug S. L., Yaszemski M. J., Bizios R., Mikos A. G. 1994. Osteoblast function on synthetic biodegradable polymers. J. Biomed. Mater. Res. 28:1445–1453. 49. Vacanti C. A., Kim W., Upton J., Mooney D., Vacanti J. P. 1995. The efficacy of periosteal cells compared to chondrocytes in the tissue engineered repair of bone defects. Tissue Eng. 1:301– 308. 50. Ishaug S. L., Payne R. G., Yaszemski M. J., Aufdemorte T. B., Bizios R., Mikos A. G. 1996. Osteoblast migration on poly(-hydroxy esters), Biotech. Bioeng. 50:443–451. 51. Pouton C. W., Akhtar S. 1996. Biosynthetic polyhydroxyalkanoates and their potential in drug delivery. Adv. Drug Del. Rev. 18:133–162. 52. Koosha F., Muller R. H., Davis S. S. 1989. Polyhydroxybutyrate as a drug carrier. Crit. Rev. Therap. Drug Carr. Sys. 6:117–130.

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16 Controlled Porosity of Tissue-Engineered Cortical Bone Grafts Kai-Uwe Lewandrowski Massachusetts General Hospital, Boston, Massachusetts Shrikar P. Bondre, Debra J. Trantolo, and Donald L. Wise Cambridge Scientific Inc., Cambridge, Massachusetts

I CLINICAL MANAGEMENT OF BONE DEFECTS The management of large skeletal defects continues to present a challenge to orthopedic surgeons, particularly when the problem arises in young patients in whom artificial devices and joint implants are likely to lead to early failure. Large frozen cortical bone allografts have been increasingly used in limb-sparing procedures not only in the treatment of bone tumors [1–3], after massive bone loss following traumatic injuries [4,5], and in the treatment of avascular necrosis [6], but also in failed joint arthroplasties [7–11], where extensive bone loss due to osteolysis is commonly encountered [12–22]. Although the overall success rate for massive cortical bone allografts, implying return to work and engagement in relatively normal activities without crutches or braces, is approximately 75–85%, only 50% of these patients have an entirely uncomplicated postoperative course. About a quarter of the total group require reoperations such as autologous grafting or replating for stress fracture delayed unions [4,23–25]. Some patients require excision of the graft because of infection [26–30], reimplantation, long-term bracing, or in some cases amputation. These results clearly indicate that problems still exist with this procedure and that if the technique is to be more widely applied, it must be more extensively examined and materially improved. Clinical and experimental reports provide sufficient evidence to conclude that massive cortical bone allografts serve primarily as osteoconductive scaffolds with poor osteoinductive potential. In fact, new bone is often found to penetrate into the graft only a 317

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few millimeters from its peripheral surfaces. In addition, studies on retrieved allografts have shown that these types of grafts seem to provoke foreign body reactions, as evidenced by the formation of a fibrous tissue cuff [31]. Resorption of the nonviable bone of allogenic cadaveric bone implants often forms the basis for clinical complications. Thus, better control of bone resorption during incorporation may ultimately improve clinical outcomes. II PREVIOUS STUDIES ON ENHANCEMENT OF INCORPORATION OF CORTICAL BONE GRAFTS Previous studies have attempted to improve host incorporation by altering the geometrical surface configuration of cortical bone [32–37]. The mechanism by which the presence of perforations promotes osteogenesis and incorporation in demineralized grafts is a function of either the greater surface area of partially demineralized bone or increased access to vascular tissue or both. These studies pointed out that the geometry of an implant may influence the extent of bone formation. For example, the osteoconductive potential of coralline hydroxyapatite with a pore size of 0.6 mm was attributed, at least in part, to its morphometric similarity to cancellous bone [38]. As such, the osteoconductive potential of cortical bone allografts treated to yield likewise structures appeared to be similar. Gendler [33] used fully demineralized diaphyseal allogeneic struts that were perforated with the use of a mechanical drill. In comparison, O’Donnell et al. [35] used demineralized calvarial cortical bone. Bernick et al. [32] characterized the inductive cellular events in a similar system. Scanlon [36] implanted demineralized canine femoral strut allografts in an orthotopic model. The latter two studies demonstrated the use of an erbium:yttrium-scandium-gallium-garnet (Er:YSGG) laser for drilling of cortical bone allografts, thereby increasing the porosity and allowing demineralization to proceed to areas that would normally be inaccessible to the demineralization process. When reimplanted, these grafts may therefore be more osteogenic than cortical grafts without holes. Although demineralization has been successful in improving osteoinductive properties of bone, as shown by the abundance of pertinent literature on experimental [39–42] and clinical studies [43–46], fully demineralized cortical bone has found limited clinical application. This appears primarily attributable to the fact that cortical bone, once subjected to extensive demineralization, has lost its essential biomechanical properties. Two recent studies [47,48] demonstrate the merit of surface modification of cortical bone allografts. A study in sheep showed that incorporation of massive cortical bone allografts may in fact be enhanced by perforation and partial demineralization. In this study, laser-perforated and partially demineralized allografts were transplanted into sheep tibia. Although perforated and partial demineralized allografts were completely incorporated at 9 months postoperatively, bone resorption originating in the laser holes and leading to widening of the hole diameter was evident as early as 2 months postoperatively. This indicated that increasing the porosity of a demineralized cortical bone allograft may not only foster the process of incorporation, but may also be accompanied by considerable bone resorption with widening of hole perforations prior to new bone formation. Engineering of polymer–allograft composites may provide a feasible adjunct to surface modifications that solely rely on the inherent osteoinductive and osteoconductive potential of cortical bone allografts.

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III CONCEPT OF POROSITY CONTROL The concept of porosity control in tissue-engineered bone constructs is based on the use of degradable polymers in conjunction with porous cortical allografts. In the PLGA family of polymers chemical hydrolysis of the hydrolytically unstable backbone is the prevailing mechanism for polymer degradation. This occurs in two phases. In the first phase, water penetrates the bulk of the polymer-filled pores, preferentially attacking the chemical bonds in the amorphous phase and converting long polymer chains into shorter water-soluble fragments. Because this occurs in the amorphous phase initially, there is a reduction in molecular weight without a loss in physical properties because the polymer matrix is still held together by the crystalline regions. The reduction in molecular weight is soon followed by a reduction in physical properties as water begins to fragment the material. In the second phase, enzymatic attack and metabolization of the fragments occur, resulting in a rapid loss of polymer mass. The rate of degradation of poly(lactic-co-glycolic acids) (PLGAs) can be controlled, in part, by the copolymer ratio with higher glycolide or lactide ratios favoring longer degradation times. Polymers of varying copolymer ratios, including poly(lactic acid) (PLA)/PLGA 75:25 and PLGA 50:50 can be used to offer extremes of degradation rates. The three polymers represent a wide range of degradation rates and nominally open the allograft pores at early (PLGA 50:50), middle (PLGA 75:25), and later (PLA) time points. A Designing a Biopolymer-Reinforced Bone Allograft Pilot studies were designed to determine if reasonable polymer–allograft composites could be prepared, reasonability being dictated by adequate adherence of the polymer to the bone in the pores of the proposed allograft. Based on varying degradation rates three types of PLGA polymers were selected. These were poly(D,L-lactic-co-glycolic acid) polymers composed of lactic acid and glycolic acid in the mole ratio of 50:50 (PLGA 50:50), in the mole ratio of 75:25 (PLGA 75:25), and in the mole ratio of 100:0 [PLA, poly(70-L-lacticco-30-D,L-lactic acid)]. 1 Sample Preparation Human cortical bone samples, obtained from a bone bank, provided diaphyseal segments of long bones. The bone was cut into slivers of 35 mm (length)  4 mm (width)  4 mm (height) using a scroll saw (Drexel). Samples were perforated with a pulsed Er:YAG laser (SEO Laser 1-2-3, Schwartz Electro Optics, Concord, MA) operating at a wavelength of 2940 nm. Holes of 330 m diameter were drilled through the entire cortex, typically generating 0.12 J in a 100- to 200-s pulse train. (A Er:YAG laser was used to drill the holes because it reduced the thermal necrosis of the allograft during the drilling.) The 35-mmlong bone samples were secured on a mount and a pattern of 20 holes in two rows were drilled along the longitudinal axis of the sample. The laser holes approximated cylinders of 330 m in diameter, with centers located 1.5 mm apart. Demineralization of bone samples was performed by immersion of samples in a well-stirred acid bath containing 25 mEq 0.5M HCl/g of bone using times obtained from mathematical models based on diffusion kinetics [49]. Based on gross variation in the degradation rates, three types of polymers from the PLGAs were selected. These were poly(D,L-lactic-co-glycolic acid) polymers composed of lactic acid and glycolic acid in the mole ratio of 50:50 (PLGA 50:50), in the mole ratio of 75:25 (PLGA 75:25), and in the mole ratio of 100:0 [PLA, poly(70-L-lactic-co-30-D,L-lac-

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tic acid)]. The polymers are commercially available from BI Chemicals as Resomer RG506 (PLGA 50:50), Resomer RG756 (PLGA 75:25), and Resomer LR 708 (PLA). Gel permeation chromatography was used to measure molecular weights calculated by comparison with polystyrene standards using a Waters instrumentation system (712 WISP autosampler, 600E system controller, and a 410 differential refractometer) equipped with 100-Å and 1000-Å Styragel columns placed in series. A THF mobile phase at a flow rate of 1.2 mL/min yielded acceptable elution times. The PLA, PLGA 50:50, and PLGA 75:25 samples were dissolved in glacial acetic acid. Vacuum techniques can be used to fill the pores of the laser-drilled cortical bone. Rectangular cortical bone constructs were added to the vial. The vial was placed in a 100mL lyophilization flask, and the polymer solution was forced into the holes using continuous evacuation and repressurization, the end point being the absence of bubbles from the laser holes. To remove the residual glacial acetic acid solvent, the polymer bone constructs were then frozen in an isopropanol–dry ice bath (79°C) and lyophilized (Labconco, Freezedryer8) at 40°C and a vacuum of less than 5 mHg for a period of 48 h. This resulted in the formation of polymer foam plugs in the laser-drilled holes. 2 In Vitro Methods The PLGA 50:50, PLGA 75:25, and PLA foam bone constructs were placed in 20-mL tubes containing 10 mL of phosphate buffered saline (pH 7.4). The tubes were then left in a water bath operating at 60 cycles per minute and 37°C. The constructs were evaluated at 1, 2, and 3 weeks. Surface scanning electron micrographs were taken at different time points to view the extent of degradation over time of the polymer from the laser-drilled holes. The images were taken on an Amray 1000 SEM. 3 Preliminary Results The observations from the scanning electron micrographs confirmed the varying degradation rates of the polymers (Fig. 1). At time 0, the polymers completely filled the drilled holes. After 1 week of incubation the PLGA 50:50 (the most quickly degrading polymer) only partially filled the drilled hole. The PLGA 75:25 at the 1-week time point showed slight degradation (as determined via surface roughening), but the polymer still substantially filled the hole. The PLA, on the other hand, showed minimal degradation; only a slight curvature was observed on the polymer surface. At the 2-week time point, the PLGA 50:50 had completely degraded and the hole was essentially empty with only slight degradation observed along the edge of the hole. At 2 weeks, the PLGA 75:25 showed partial degradation; the hole was partially filled with polymer. In comparison, the PLA still filled the hole with only slight degradation observed at the edge of the hole. This pilot study was useful for two reasons: (1) early methods for preparing the polymer/bone composites were developed and (2) the differing degradation behaviors were confirmed. B Design Considerations The poly(lactic-co-glycolic acids) are a family of resorbable polymers comprised of lactic and glycolic acid copolymers. The rate of degradation of PLGAs can be controlled, in part, by the copolymer ratio, with higher glycolide or lactide ratios favoring longer degradation times. The disproportionate bone resorption in highly porous allografts may be circumvented by filling the graft perforations with resorbable PLGAs that degrade over time.

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Figure 1 Observations from scanning electron micrographs.

As such, the filled allografts can program the developing porosity while maintaining satisfactory mechanical strength during incorporation. The net result of this process would be improved replacement of the allogenic bone by the host’s own bone without eminent risk for premature mechanical failure due to overstimulated loss of bone by bone allograft resorption. C Experimental Outcome Parameters Values are assigned to the in vitro mechanical outcomes, which directly relate to polymer degradation in the perforations. Ideally, the polymer-modified graft provides for rates of polymer degradation commensurate with the process of bone allograft incorporation. Clinically applied laser perforation and demineralization result in a loss of approximately 20% in the overall mechanical properties of the graft.

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The first issue to be addressed is the timing of this available “released” porosity. In addressing this, one can invoke the stages of bone induction within the surrounding host bed and then apply the analogy of Vandersteenhoven and Spector [42]: Stage of induction of mesechymal cells of the host bed: Stage of fibrous and vascular tissue invasion: Stage of de nove bone formation in tissues surrounding the graft: Stage of increasing incorporation by bone resorption and new bone formation:

Up to about 1 week To ~2–4 weeks To ~4–6 weeks Several months to years

When bone formation at the junction sites is complete—average of about 6–12 months—the graft is considered to be clinically and radiographically healed. Once new bone is formed at the graft site, immature bone starts to respond to stresses by remodeling to highly structured bone, as indicated by formation of lamellar structures (rodents) or secondary osteons (humans). Depending on the site of transplantation; quality of the host bed; means of fixation; and other concurrent factors such as the presence of infection, chemotherapy, or radiation, successful incorporation is evident both clinically and radiographically with the majority of the bone graft being incorporated in about 8–12 months with about 50% recovery of mechanical properties in about 3 weeks [50]. Given the biology of the incorporation process of large cortical bone allografts, initial evaluation as early as 3 weeks and long-term evaluation appear as acceptable targets for timing of the mechanical support of the polymer filling. The second concern to be addressed is the potential of the polymer filling to augment the initial mechanical properties such that mechanical support is feasible. Here it is instructive to compare the flexural properties of some biodegradable PLGA polymers with similar values for bone as given in Table 1 [51] and Table 2 [52]. Under ideal conditions, the rate of mechanical losses in the composite should equal the rate of mechanical recovery in the bone throughput over the course of bone healing. From the data shown in these tables, it appears feasible to suggest that the initial polymer filling should augment the flexural properties of the demineralized/laser-perforated graft. Thus, using these results and as-

Table 1 Flexural Properties of Biodegradable Polymers Polymer

Flexural strength, MPa

Flexural modulus, GPa

PLLA PLLA PLLA/PDLLA (90:10) PLLA/PDLLA (75:25) PLLA/PDLLA (50:50) PLLA PLLA PGA/PLA (2:98) PGA/PLA (90:10)

109 129.7 (113–142) 129.5 (113–142) 115 97 67.7 (45–83) 132 (119–145) 65 95 (40–150)

6.27 (2.4–10.3)

4.7 (4–5)

Notes: PLLA  poly(L-lactide); PDLLA  poly(D,L-lactide); PGA  poly(glycolide); PGA/PLA  poly(glycolide-co-lactide). Average value given with range in parentheses. Source: Ref. 51.

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Table 2 Flexural Properties of Bone Bone Human femur Monkey tibia Cattle femur Sheep metatarpus Horse metatarpus, metacarpus Rabbit femur

Flexural strength, MPa

Flexural modulus, GPa

103–238

9.1–15.7 9.0  1.3 18.1  19.4 18.9  2.2 14–18.4

209–228 195–240 130  5

Source: Ref. 52.

signing values to the initial composite mechanics and to the rate of mechanical loss in the early course of the in vitro experiments, one should be able to assign early in vitro success criteria on the basis of an overall limitation of mechanical strength to no more than 25–30% at any given time during the incorporation process. IV DISCUSSION The engineered cortical bone transplants are expected to incorporate faster, having the potential to develop increasing porosity over time. The approach of using biopolymers of varying biodegradation rates for filling of graft perforations is based on the large body of “slow release” literature, which provides sufficient evidence that varying polymer degradation rates can be exploited for specific biomedical applications while at the same time permitting use of the polymers as carriers of active molecules or factors or even as reinforcing elements. The use of biopolymer fillings in perforations of partially demineralized bone should result in an overall enhancement of incorporation for two reasons. First, the outer demineralized surface layer provides a highly osteoinductive matrix for the ingrowth of host mesenchymal cells from the adjacent soft tissues into the allograft that can be transformed in bone-forming chondrogenic and osteogenic cells. Second, it may further promote the osteoinductive concept by the development of focal centers of new bone formation in patent holes due to the release of growth factors from the demineralized bone matrix into the osteoconductive polymer scaffolds. Future in vitro and in vivo studies can be undertaken to determine which PLGA polymers have the appropriate biodegradation rates suitable for both bonding with cortical bone and maintenance of sufficient hole closure for the desired interval until patency develops. An in vivo study using rats is planned for comparative screening of candidate composite systems by performing mechanical, histologic, and histomorphological analyses of graft incorporation. Additional outcome variables will include (1) in vitro and in vivo mechanical properties of polymer–allograft composites, (2) extent of incorporation, and (3) rate of incorporation. Feasibility of the proposed approach entails the ability to control allograft resorption and subsequent bone cell ingrowth into perforations by osteoconductive polymer fillings and their ability to modulate the extent of bone resorption in light of an overall increased graft porosity. However, limitation of mechanical strength with increasing graft porosity due to progressing polymer degradation appears crucial and should not exceed 25–30%. Previous in vivo studies with perforated and partially demineralized cortical bone

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allografts indicate new bone ingrowth in perforations to be between 3 to 4 weeks [47]. Therefore, an initial porosity equaling a loss of mechanical strength of no more than 25–30% should be present by the graft to the surrounding soft tissue host bed. ACKNOWLEDGMENT The authors acknowledge the partial support of this work under Grant No. 1 R43 AR46608 (to D. J. Trantolo) from the National Institutes of Health/National Institute of Arthritis and Musculoskeletal and Skin Diseases. REFERENCES 1. Clohisy D. R., Mankin H. J. 1994. Osteoarticular allografts for reconstruction after resection of a musculoskeletal tumor in the proximal end of the tibia. J. Bone Joint Surg. Am. 76:549–554. 2. Mankin H. J., Gebhardt M. C., Jennings L. C., Springfield D. S., Tomford W. W. 1996. Longterm result of allograft replacement in the management of bone tumors. Clin. Orthop. 324:86–97. 3. Rougraff B. T., Simon M. A., Kneisl J. S., Greenberg D. B., Mankin H. J. 1994. Limb salvage compared with amputation for osteosarcoma of the distal end of the femur. A long-term oncological, functional, and quality-of-life study. J. Bone Joint Surg. Am. 76:649–656. 4. Jaffe K. A., Morris S. G., Sorrell R. G., Gebhardt M. C., Mankin H. J. 1991. Massive bone allografts for traumatic skeletal defects. South. Med. J. 84:975–982. 5. Mahomed M. N., Beaver R. J., Gross A. E. 1992. The long-term success of fresh, small fragment osteochondral allografts used for intraarticular post-traumatic defects in the knee joint. Orthopedics 15:1191–1199. 6. Flynn J. M., Springfield D. S., Mankin H. J. 1994. Osteoarticular allografts to treat distal femoral osteonecrosis. Clin. Orthop. 303:38–43. 7. Convery F. R., Minteer-Convery M., Devine S. D., Meyers M. H. 1990. Acetabular augmentation in primary and revision total hip arthroplasty with cementless prostheses. Clin. Orthop. 167–175. 8. Gross A. E., Allen G., Lavoie G. 1993. Revision arthroplasty using allograft bone. Instr. Course. Lect. 42:363–380. 9. Huo M. H., Friedlaender G. E., Salvati E. A. 1992. Bone graft and total hip arthroplasty. A review. J. Arthroplasty 7:109–120. 10. Kraay M. J., Goldberg V. M., Figgie M. P., Figgie H. E. 1992. Distal femoral replacement with allograft/prosthetic reconstruction for treatment of supracondylar fractures in patients with total knee arthroplasty. J. Arthroplasty 7:7–16. 11. Pak J. H., Paprosky W. G., Jablonsky W. S., Lawrence J. M. 1993. Femoral strut allografts in cementless revision total hip arthroplasty. Clin. Orthop. 172–178. 12. Bauer T. W., Muschler G. F. 2000. Bone graft materials: an overview of the basic science. Clin. Orthop. Rel. Res. 371:10–27. 13. Cohen D. B., Chotivichit A., Fujita T., Wong T. H., Huckell C. B., Sieber A. N., Kostuik J. P., Lawson C. H. 2000. Pseudarthrosis repair: autogenous iliac crest versus femoral ring allograft. Clin. Orthop. Rel. Res. 371:46–55. 14. Ehrler D. M., Vaccaro R. A. 2000. The use of allograft bone in lumbar spine surgery. Clin. Orthop. Rel. Res. 371:38–45. 15. Haddad F. S., Spangehl M. C., Mark J., Masri B. A., Garbuz D. S., Duncan C. P. 2000. Circumferential allograft replacement of the proximal femur: a critical analysis. Clin. Orthop. Rel. Res. 371:98–107. 16. Head W. C., Malinin T. I. 2000. Results of onlay allografts. Clin. Orthop. Rel. Res. 371:108– 112.

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17 Biodegradable Scaffolds as Bone Graft Extender Kai-Uwe Lewandrowski Massachusetts General Hospital, Boston, Massachusetts Shrikar P. Bondre, Debra J. Trantolo, and Donald L. Wise Cambridge Scientific, Inc., Cambridge, Massachusetts

I CLINICAL PROBLEM According to a survey by Medical Data International, Inc. (1999) on surgical procedural volumes in the United States, bone grafts (both autografts and allografts) were required in over 425,000 musculoskeletal operations performed in 1995. Overall, 80% of bone grafts were used in orthopedic and spine surgery. The primary indications for bone grafts in orthopedics were for fracture repair, defect filling, and total joint revisions, and in spine surgery for fusion procedures. Arguably, the most important criterion used by orthopedic and spine surgeons in assessing a bone replacement technology is its healing potential. The products that are currently available on the market are osteoconductive in nature and, although FDA studies have shown that these products produce comparable results to autogenous bone, most surgeons have been reluctant to use them, particularly in nonapproved indications [1]. It was estimated that in 1995 synthetic bone replacement materials were used in only 25,000–30,000 of the clinical grafting procedures. A biodegradable bone graft extender with inductive capacity could expand the clinical utility of bone replacement materials. Inductive biodegradable bone graft extenders could prove extremely useful in situations where use of autografts is preferable yet the patient’s own bone stocks may be insufficient. In addition, such materials could pose a significant advantage over allografts, which work inconsistently in a variety of surgical procedures, by eliminating the slight but troubling risk of introducing a variety of viruses, including those that cause AIDS or hepatitis. Further, a biopolymeric extender that has 327

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scaffold properties could provide a framework into which host bone can regenerate and heal. Bone cells could weave into and through the porous microstructure of a biodegradable inductive bone graft extender, thereby providing a regenerative structure to support the new tissue, blood cells, and soft tissue as they grow to connect fractured or defective bone segments. If successful, the use of a porous scaffoldlike biodegradable material which extends, i.e., increases the volume of, autologous bone to fill skeletal defects in reconstructive or fusion procedures would be extremely safe and effective. II CLINICAL MANAGEMENT OF BONE DEFECTS The management of skeletal defects has continued to present a challenge to orthopedic surgeons, particularly when larger amounts of autologous bone graft material are needed. This problem commonly arises in revision surgeries where bony defects are common and autologous bone stocks may have been depleted during prior surgeries. Autologous bone grafting has been considered the gold standard of bone transplantation with superior biological outcomes [4]. Allografts have been used as alternative materials. Although the use of allografts eliminates the need for a second surgery, the grafted bone may be incompatible with the host bone and ultimately be rejected. Allografts also pose a slight but still troubling risk of introducing a variety of viruses including those that cause AIDS or hepatitis in the patient. In addition, allografts have proven to work inconsistently in a variety of surgical procedures. The following are clinical examples where there has been a need for bony reconstruction with a void-filling graft material which behaves like cancellous autograft. Particularly, defects large enough to render themselves unsuitable for traditional autologous bone grafting—such as the reconstruction of defects caused by surgical debridement of infections, previous surgery, tumor removal, trauma, implant revisions, and for joint or spinal fusions—could benefit from an autograft extender. The treatment of chronic bone infections usually requires, in addition to antimicrobial therapy, surgical debridement to remove necrotic tissue. Once the infection has been adequately treated, reconstruction of the surgical defect is dependent on the actual size of the defect. Autologous bone grafting is preferable; however, the amount of autologous bone that can be retrieved from the patient may not be adequate. In addition, if the patient has had previous revisions of the infectious site including previous bone grafting, autologous bone stocks may have been depleted. A bioresorbable bone graft extender that is compatible with autologous bone graft materials could be applied to larger defects. This composite could have biological properties that are superior to the cortico-cancellous allogeneic bone chips routinely used in this situation, because highly osteoinductive autograft material would be used in conjunction with a highly osteoconductive scaffold suitable for both induction and expansion of new bone formation. Another potential application of a composite graft extender could be in the reconstruction of osteolytic defects. Biologically inert methyl methacrylate–based bone cements have been commonly employed in the treatment of osteolysis in those patients where revision of a total hip arthroplasty had become necessary. Revision total hip replacements are more complex than primary total hip replacements because following the removal of the old implant and the old cement, and before cutting the new canal cavity, it is necessary to fill or pack skeletal defects in the acetabulum so that the new implant socket is held securely place. A bioresorbable poly(propylene glycol-co-fumaric acid) (PPF)–based bone graft extender which has mechanical properties similar to those of bone and supports bony ingrowth from the defect site could prove extremely useful in

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eliminating long-term complications related to the use of allograft or permanent methyl methacrylate–based bone cements. III CURRENTLY USED BONE GRAFT SUBSTITUTES AND EXTENDERS Although several companies have new bone replacement materials (BRM) products in the FDA pipeline, four companies currently compete in this market. Three of these firms, Interpore International (Irvine, CA), Collagen Corporation (Palo Alto, CA) in collaboration with Zimmer (Warsaw, IN), and Wright Medical (Memphis, TN) have products approved for fracture indications. Wright recently received 510(k) approval for a calcium sulfate BRM called OsteoSet by claiming substantial equivalence to a product previously marketed by Johnson & Johnson (New Brunswick, NJ) before enactment of medical device regulations in 1976. Currently, the leading contender in BRM (50% of the total market) is allograft processor Osteotech (Shrewsbury, NJ), with their product Grafton, which consists of allogeneic demineralized bone matrix (DBM) combined with glycerol. Available in a gel form and more recently as a putty and a flexible strip, this material is often used in conjunction with a bone graft and functions more as an extender than as a bone graft substitute. Because it is derived from human banked bone, it is hypothesized that Grafton contains bone morphogenic proteins that impart an inductive capacity to a bone graft and enhances its propensity to heal; however, this has yet to be proven. The processing methods used to guarantee sterility and viral inactivation of allogeneic tissue may also impair its osteoinductive properties. Another frequently used allogeneic bone graft extender is cortico-cancellous bone chips. Unfortunately, with this product as well, firm evidence of osteoinduction with the use of allograft extenders has not been forthcoming. Indeed, as shown in Fig. 1, allograft chips have proven to be primarily osteoconductive. In this example, allograft chips were mixed with cancellous autologous bone and retrieved postoperatively after grafting in a failed posterior spinal fusion. As shown by the arrows, there is fibrous tissue invasion of the allograft with only occasional bone ingrowth. These images have demonstrated the poor

Figure 1 Allograft chips mixed with cancellous autlogous bone retrieved 6 months postoperatively after a grafting in a failed posterior spinal fusion.

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osteoinductive characteristics of allograft extenders. If only osteoconductive, the foregoing evidence argues for the use of a biodegradable osteoconductive extender that would resorb to accommodate reconstruction and new bone formation. The bone graft substitutes and extender materials were by Interpore International. These include bone biologics products which are comprised of coralline-based bone graft substitute products, hydroxyapatite products, and other products for collecting growth factors for the express purpose of encouraging bone growth. Bone graft substitutes such as these and OEM hydroxyapatite products are derived from coral using a proprietary manufacturing process. Under this process, the exoskeleton of marine coral, which is made of calcium carbonate, is either partially or fully converted to calcium phosphate. These biocompatible materials have interconnected porosity, architecture, and chemical composition similar to that of human bone, and facilitate both bone and tissue ingrowth. When blocks or granules of these products are placed in a bone defect, they have been proven to provide a lattice for the patient’s own bone to grow through and heal. Bone biologics products include Pro Osteon 500 Porous Bone Graft Substitute (Pro Osteon 500), Pro Osteon 200 Porous Bone Graft Substitute (Pro Osteon 200), Pro Osteon 500R Resorbable Bone Graft Substitute (Pro Osteon 500R), Interpore 200® Porous Hydroxyapatite Bone Void Filler (Interpore 200), orbital implants and AGF™ (Autologous Growth Factors™). The last of these products employs concentrated growth factors derived from platelets taken from a patient’s own blood, which are then used to encourage more complete and rapid bone growth in bone grafts. Interpore is the exclusive licensee, in the areas of bone and cartilage, of a technology that produces and concentrates these growth factors. In this process, blood from the patient is first separated into different component layers using a device called a cell saver, which is routinely available in the orthopedic surgery suite. The layer containing platelet-rich plasma is then processed through a proprietary filtering technology which superconcentrates the platelets and fibrinogen. The platelets release platelet-derived growth factor (PDGF) and transforming growth factor-beta (TGF-). The fibrinogen converts to fibrin, which is generated as a gel-like material with favorable handling characteristics. The AGF can then be combined with a bone graft material such as Pro Osteon, for example, and implanted at the site of the bone deficit. Autologous Growth Factor has been shown in animal studies to increase the amount of bone growth in defects. Interpore believes that AGF provides the surgeon with the growth factors desired for faster and more complete bone graft healing and that AGF may be safer and more economical than synthetic growth factors being developed by other companies. In order to expand the clinical utility of BRMs, manufacturers are targeting other areas, e.g., spinal applications, whereas biotech firms tend to focus on developing inductive products. New materials target improved handling characteristics (e.g., injectability, moldability, and cure rates for those products that harden in situ) as well as mechanical and inductive properties. Because cost will be the determining factor in how these products are ultimately used, whether on a routine or on a selective basis, the challenge will be for manufacturers to demonstrate how these products improve patient outcomes and lower overall cost by either eliminating a need for hardware or expediting patient function. IV A BIODEGRADABLE BONE GRAFT EXTENDER Preliminary investigations looked at the feasibility of enhancing the regeneration of skeletal tissues with an extender fashioned from the biopolymer poly (propylene glycol-co-fumaric acid). This unsaturated polymer can be crosslinked with a vinyl monomer, vinyl

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pyrrolidone (VP), in the presence of effervescent agents, sodium bicarbonate (SB) and citric acid (CA), and the filler hydroxyapatite (HA) to create an osteoconductive foaming network. The biopolymer can be mixed with the autograft immediately prior to defect filling. The polymer/graft composite expands and cures in situ to a hard dimensionally stable mass having intimate contact with the defect interfaces. This work has had as its overall objective the development of a new type of bone replacement material with both improved osteoconductive and osteoinductive properties by promoting bony ingrowth into a porous scaffoldlike extender material. It is hypothesized that resorbability would support complete healing of the bone defect site and eliminate long-term complications observed with other permanent implant materials. Ultimately, this type of bone graft extender would also be amenable to tissue engineering applications by delivering growth factors and other bone-forming factors or cells via controlled release from the biopolymeric component. In short, the bone graft extender would be incorporated via the following clinical scenario: (1) the procurement of cancellous autograft bone from the patient undergoing surgical repair, (2) the mixing of the patient’s autologous material with the osteoconductive poly(propylene fumarate) scaffold, and (3) the grouting and in situ curing of the resulting extended autologous bone graft into the patient’s bone defect. The foam embodiment of a PPF-based grout has been evaluated in both in vitro and in vivo environments in an attempt to qualify its scaffold properties. In addition, it has been tested in vivo in a rat tibial defect in conjunction with various types of clinically used cancellous bone grafts employing the well-utilized rat tibial defect model of Gerhart et al. [2]. Preliminary results have suggested that PPF bone graft extenders have the ability to sustain bone ingrowth over time. In addition, the dimensional stability of the scaffold has been found to preserve the integrity of the defect site [3]. Preliminary qualitative histologic and histomorphometric studies presented herein have shown evidence of more vigorous new bone formation and migration of bone-forming cells into the PPF scaffold when mixed with autologous bone graft material. Migration had occurred as early as 3 days after implantation, which suggests a material highly suitable for stimulation of bone cell ingrowth. V ENGINEERING OF BIODEGRADABLE BONE GRAFT EXTENDERS The appeal of using a bioresorbable scaffold having an open window structure as an expanding carrier for autologous bone graft material stems from the fact that bone has a considerable potential for regeneration. In fact, bone has been considered by some to be the prototypical model for tissue engineering [5,6]. Development of composites of cells responding to regulatory signals has been at the center of engineering skeletal tissues. In fact, leading strategies on engineering and regeneration of skeletal tissues have been largely based on observations on local bone induction after implantation of bony matrices into subcutaneous sites. These models have allowed the study of sequential limb morphogenesis and have permitted the isolation of bone morphogens, such as bone morphogenetic proteins (BMPs), from demineralized adult bone matrix. The BMPs initiate, promote, and maintain chondrogenesis and osteogenesis. It appears reasonable to consider a biodegradable bone graft extender as a designable scaffold that will function as a cell and tissue carrier emphasizing material superstructuring. Therefore, current investigations have been aimed at combining autologous cancellous bone with an engineered, biodegradable superstructure that could provide optimal spatial and nutritional conditions for migration of bone cells by maintenance of the arrangement

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of structural elements (e.g., pores) for cell-to-cell contact [7]. Predictable three-dimensional matrices for engineering of bone replacement materials are still needed [8,9]. Therefore, optimization has been centered around the ability to construct a predictably and dimensionally stable scaffold bone graft extender that maintains a three-dimensional superstructure. VI POLY(PROPYLENE GLYCOL-CO-FUMARIC ACID) AS A MODEL BIOPOLYMER In the design of a biodegradable scaffold for use as a bone graft extender, it is important to consider the function of the scaffold as well as the normal healing process of bone defects. An immediate function of a bone graft extender should be osteoconduction, that is, provision of a framework or scaffolding for vascular induction, cellular infiltration and attachment, cartilage formation, and calcified tissue deposition. This function requires initial maintenance of porosity after placement of the graft. Simultaneously, the proposed bone graft extender would have significant osteoinductive potential based on its incorporation of autologous cancellous bone. A In Vitro Studies A dimensionally stable porous scaffold based on the crosslinking of the unsaturated PPF polymer with a vinyl monomer, vinyl pyrrolidone, in the presence of the effervescent reagents sodium bicarbonate and citric acid, and an osteoconductive filler, hydroxyapatite, has been considered as a potential bone graft extender. To demonstrate feasibility, the PPF scaffold was evaluated in vitro for morphological, mechanical, and surface properties and in vivo using a rat tibial defect model established by Gerhart et al. [2]. Scaffolds were found to be mechanically comparable to trabecular bone, dimensionally stable, and porous. Histologic and histomorphologic examination of the implant region of the rat tibial defect suggested that the scaffold of the biodegradable bone graft extender supports bony ingrowth and the stability of the scaffold preserves the dimensional integrity of the defect site. The bone graft extender carrier is foaming putty which has been prepared as a two part system in order to separate the polymerizable entities (PPF, VP) from the initiator benzoyl peroxide (BP). The foaming agent consisted of a stoichiometric ratio of citric acid and sodium bicarbonate (i.e., 1.0:1.3, w/w). A typical formulation has been included in Table 1. Table 1 Typical Formulation of the PPF Bone Graft Extender Wt% PPF Hydroxyapatite Vinyl pyrrolidone Ethanol Sodium bicarbonate Benzoyl peroxide Citric acid Water

39.32 11.38 12.10 12.70 1.95 1.71 1.47 19.37

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Figure 2 Scanning electron micrograph of PPF bone graft extender carrier at t  0 (20).

B Mechanical Characterization of the Bone Graft Extender Mechanical testing has been performed in accordance with ASTM F451-95 for bone cements. Formulations were cured in 0.25-in. diameter by 0.50 in. high cylindrical Teflon molds. These cured samples were compression tested on an Instron Model 8511 materials testing machine fitted with a 500-lb load cell. The samples were compressed to failure at a crosshead speed of 1 cm/min and the load deformation curve was recorded. From these data the ultimate compressive stress and Young’s modulus could be calculated. The ultimate compressive stress was found as the applied load at failure divided by the original cross-sectional area of the test specimen, and the Young’s modulus as the slope of the load deformation curve in its linear portion. The mean compressive strength of the porous cement was determined to be 6.77  2.5 MPa (n  8). These compressive strength data were comparable with values that have been reported by Carter and Hayes, having measured this property of bone as a function of strain rate and density. At a comparable strain rate, the compressive strength of trabecular bone (  0.31 g/cm3) was noted to be 5.0 MPa [10]. C Morphological Characterization of the Bone Graft Extender Given osteoblast migration into the graft extender as a design goal, scaffold porosity was desirable. It has been determined that this PPF bone graft extender carrier is characterized by small and large pores. A scanning electron micrograph showing the presence of both types of pores is shown in Fig. 2. The scaffold has a wide pore size distribution (median pore size: 67 m) with at least 30% of pores having an average diameter greater than 200 m (as confirmed by mercury intrusion porosimetry). After 24 days in vitro (PBS, pH 7.2, 37°C), no gross change was seen in the dimensions of the implants, and scanning electron microscopy confirmed maintenance of the trabecular-like structure (Fig. 3). The porosity was tunable by controlling the citric acid content, as shown in Table 2.

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Figure 3 Scanning electron micrograph of PPF bone graft extender after 24 days in vitro (PBS, pH 7.23, 37°C) (50).

D Osteoconductivity and Biocompatibility of PPF Foams To evaluate the suitability of a bioresorbable PPF-based foam for bony ingrowth, an initial in vivo study in rats was conducted using the tibial defect model of Gerhart et al. [2]. The scaffold formulation was mixed with either autograft or allograft material immediately prior to placement directly into a cylindrical metaphyseal defect made into the anterior aspect of the rat tibia using a dental cutter (measuring 4.5 mm in diameter). The experimental design also contained control formulations consisting of defects without any graft material, autografts, and PPF alone. Groups of eight animals were sacrificed at 1 and 4 weeks postoperatively in order to investigate early stage biocompatibility. Histomorphometry was performed based on a minimum of eight serial longitudinal sections that averaged the amount of new bone formation over the entire implant area, thereby allowing bone ingrowth to be expressed as a percentage rate for direct comparison between experimental groups. Histologic evaluation of negative controls, or the unfilled tibial defects, found the defect to be filled in with fibrous and granulation tissue at 1 week postoperatively. At 4 weeks, evidence of some reactive new bone formation was seen without complete closure of the defect (Fig. 4a,b). Defects that were filled with either PPF or autograft bone ma-

Table 2 Influence of Citric Acid on the Pore Size of PPF Scaffolds Citric acid (mL)

Pore size (m)

1.06 0.86 0.63

67 29 0.0123

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Figure 4 Photomicrographs of longitudinal sections (H&E) of rat tibia. (a) Empty drill hole defect at 1 week. (b) Empty drill hole defect at 4 weeks. There is lack of healing of the empty drill hole defect at 1 week with some reactive bone formation at 4 weeks. (c) Autograft at 1 week. (d) Autograft at 4 weeks. There is extensive new woven bone formation in the intramedullary space under the autologous bone graft at 1 week and almost complete healing at 4 weeks. (e) PPF foaming extender alone at 1 week. (f) PPF foaming extender alone at 4 weeks. Implantation of the PPF foaming extender resulted in some early bone ongrowth at 1 week. This new bone formation appeared to have closely followed the shrinking and degrading implant at 4 weeks.

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terial alone were used as positive controls. These sections showed early new woven bone formation in the autograft group with near complete healing of the defect at 4 weeks (Fig. 4c,d). The defects containing the PPF alone were found to have early new bone formation at 1 week postoperatively. Evidence of this formation having been promoted by the PPF foaming scaffold demonstrates the osteoconductive properties of the material. Increasing resorption of the material with subsequent gradual ingrowth of new bone was seen at 4 weeks postoperatively (Fig. 4e,f). The combination of the PPF bone graft extender with either allograft or autograft material has resulted in enhancement of new bone formation with both the allo- and autograft groups (Fig. 5a–d). However, this improved osteoinduction was only seen when the PPF bone graft extender was mixed with fresh autograft. The combination of PPF with autograft has resulted in more new bone formation than when just the autograft was used alone. In addition, an increase in new bone formation was found when a combination of allograft and PPF was used versus allograft by itself. It is hypothesized that this increase was due to the additional osteoconductive ef-

Figure 5 Photomicrographs of longitudinal sections (H&E) of rat tibia. (a) Allograft  PPF foaming extender at 1 week. (b) Allograft  PPF foaming extender at 4 weeks. There is new bone formation surrounding the graft at 1 week which appeared more extensive at 4 weeks. (c) Autograft  PPF foaming extender at 1 week. (d) Autograft  PPF foaming extender at 4 weeks. Implantation of the PPF foaming extender resulted in extensive early new bone formation in the proximity of the autograft at 1 week. At 4 weeks, there is complete closure of the defect. Most of the polymeric PPF-based extender material is resorbed.

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fect of the PPF. These histologic observations were supported by histomorphometric measurements of new bone formation that demonstrated significant increases when PPF was used in combination with either auto- or allograft. In a similar rodent study, the temporal sequence of bone ingrowth into a PPF foaming scaffold injected into tibial drill holes was investigated for osteoclastic and osteoblastic activity, and neovascularization was found in the foam implantation site as early as 1 week postoperatively [3]. At postoperative week 3, the drill hole was completely healed in all animals injected with the foam. In comparison, this was not the case in control animals that had not had an implant injected into the defect. Although there was some periosteal bone formation at 3 and at 4 weeks postoperatively, complete healing of the hole had not been observed in any of the sham operated animals. VII CONCLUSIONS It has been determined that increasing the volume of autologous bone graft material with a biodegradable porous bone graft extender for applications for regeneration of skeletal tissues could, in fact, be feasible. The biopolymeric extender for the bone grafts was fashioned from poly(propylene glycol-co-fumaric acid). The overall objective of these studies has been the development of a new type of biocompatible bone graft extender with improved osteoconductive and osteoinductive properties that will not compromise clinical outcomes by the use of allogeneic or other suboptimal bone graft materials. The PPF bone graft extender has been found to be morphologically similar and biologically compatible with autologous cancellous bone. It remains to be shown if this will also prove to be true when used in much larger quantities as needed for the reconstruction of significant skeletal defects. Continued work focuses on preparing a polymeric scaffold carrier as a precursor to the developing new bone, which will ultimately stem from the autologous cancellous bone mixed within the carrier and from migration of bone cells into the scaffold from the surrounding host bed. The use of autologous cancellous bone graft within the biodegradable bone extender carrier will conceivably enhance new bone formation at the repair site while eliminating long-term complications related to the use of the carrier material because of its resorbability. Analysis of mixings containing both autologous cancellous bone with the curing biodegradable bone graft extender and development of material combinations that demonstrate in vitro and in vivo durability over time, while supporting bone formation equivalent to traditional autologous bone grafting, will be necessary prior to clinical application. However, if these properties could be demonstrated, a biodegradable porous bone graft extender could be readily applied for increasing the volume of autologous cancellous bone to numerous clinical repair scenarios due to its greater overall morphologic similarity to cancellous bone. In addition, such bone graft extender composites could be further utilized by seeding with bone growth factors or bone-forming cells to improve osteoinductive and osteoconductive properties. ACKNOWLEDGMENT The authors acknowledge the partial support of this work under Grant No. 1 R43 AR46970 (to D. J. Trantolo) from the National Institutes of Health/National Institute of Arthritis and Musculoskeletal and skin Dieases.

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REFERENCES 1. Health Advances, Inc. 1999. U.S. market for bone graft substitutes [proprietary]. Health Advances, Inc.: Wellesley, MA. 2. Gerhart T. N., Renshaw A. A., Miller R. L., Noecker R. J., Hayes W. C. 1989. In Vivo histologic and biomechanical characterization of a biodegradable particulate composite bone cement, J. Biomed. Mater. Res. 23:1–16. 3. Lewandrowski K.-U., Cattaneo M. V., Gresser J. D., Wise D. L., White R. L., Bonassar L., Trantolo D. J. 1999. Effect of a poly(propylene fumarate) foaming cement on the healing of bone defects. Tissue Eng. 5(4):305–316. 4. Gadzdag R. A., Lane J. M., Glaser D., Forster R. A. 1995. Alternatives to autogenous bone graft: efficiency and indications. J. Am. Acad. Orthop. Surg. 3(1):1–8. 5. Laurencin C. T., Attawia M. A., Elgendy H. E., Herbert K. M. 1996. Tissue engineered bone regeneration using degradable polymers: the formation of mineralized matrices. Bone 19(Supp. 1): 93S–99S. 6. Reddi A. H. 1998. Role of morphogenetic proteins in skeletal tissue engineering and regeneration. Nat. Biotechnol. 16(3):247–252. 7. Wintermantel E., Mayer J., Blum J., Eckert K. L., Luscher P., Mathey M. 1996. Tissue engineering scaffolds using superstructures. Biomaterials 17(2):83–91. 8. Puelacher W. C., Vacanti J. P., Ferraro N. F., Schloo B., Vacanti C. A. 1996. Femoral shaft reconstruction using tissue-engineered growth of bone. Int. J. Oral Maxillofac. Surg. 25(3):223– 228. 9. Rivard C. H., Chaput C., Rhalmi S., Selmani A. 1996. Bio-absorbable synthetic polyesters and tissue regeneration. A study of three-dimensional proliferation of ovine chondrocytes and osteoblasts. Annales de Chirurgie 50(8):651–658. 10. Carter D. R., Hayes W. C. 1976. Bone compressive strength: the influence of density and strain rate. Science 194:1174–1175.

18 Bioactivity of Nanohydroxyapatite in a Scaffold for Periodontal Repair Kai-Uwe Lewandrowski Massachusetts General Hospital, Boston, Massachusetts Shrikar P. Bondre, Donald L. Wise, and Debra J. Trantolo Cambridge Scientific, Inc., Cambridge, Massachusetts Jacqueline Y. Ying Massachusetts Institute of Technology, Cambridge, Massachusetts

I NANOPARTICLE TECHNOLOGY FOR BIOPOLYMERIC SCAFFOLDS Nanoparticle technology has been proposed for the fabrication of stronger bioactive biopolymeric bone graft materials. Nanoparticulate hydroxyapatite (nano-HA) has been considered as a means for improving the ability of biopolymeric scaffolds to support orthopedic and dental implants while consolidation of bone grafts occurs. In addition, nanohydroxyapatite employed for this particular application may improve upon the biodegradable scaffold’s mechanical and molding properties. This in turn should enhance long-term maintenance of implant contouring and should contribute to the overall longevity of the repair site. Despite demonstrated improvement in mechanical and molding properties, relatively little is known about the bioactivity of nanoparticulate hydroxyapatite–augmented biomaterials when placed in bony defects. Newer bone replacement materials are currently employed for alveolar and mandibular ridge reconstruction. Other techniques include the use of biodegradable membranes for guided periodontal tissue regeneration during bony recovery after grafting procedures. Advanced polymer systems are in the process of being developed to accommodate biomaterial needs in maxillofacial and dental bone reconstruction. However, despite significant advances, it has remained a chal339

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lenge to develop clinically applicable bone replacement materials. This problem has been, at least in part, due to the difficulty of producing sufficient bony ingrowth for prolonged periods of time thereby affecting implant fixation. Thus, bony ingrowth into such materials in skeletal repair sites is often limited to the periphery of the implant rather than a deep tissue penetration. The latter process, however, appears eminently important for the successful development and manufacturing of viable tissue equivalents. The inability to develop biopolymeric scaffolds that better support penetration with skeletal cells and tissue substrates has somewhat limited broader application of these materials to clinical practice. In an attempt to further enhance the bioactivity of biopolymeric-based scaffolds for dental and maxillofacial applications, the augmentation of biodegradable scaffolds with nanoparticulate hydroxyapatite has been suggested. Biodegradable and biocompatible polymeric scaffolds have been used as orthopedic biomaterials for the promotion of tissue regeneration [1,3]. Degradable biopolymers have been shown to facilitate bony ingrowth [4] and have served as excellent substrates for cell and tissue culture [5,6]. Hydroxyapatite has been blended with different biopolymers to improve mechanical properties as well as to increase the osteconductive nature of the composite. Hydroxyapatite has proven to be an attractive and widely utilized bioceramic material for orthopedic and dental implants because it closely resembles native tooth and bone crystal structure. Other common biopolymers that have been considered as tissue engineering scaffolds for skeletal repair include those derived from Kreb’s cycle monomers, such as the polylactic acids and polyglycolic acids. Another related biopolymer commonly considered for this type of application is the unsaturated polyester, poly(propylene glycol-cofumaric acid) (PPF). The PPF copolymer was chosen as the model biopolymer scaffold to investigate the feasibility of inclusion of hydroxyapatite nanoparticles into a biopolymer system. The ideal feature of PPF is its unsaturation; PPF can be crosslinked (“cured”) to yield a material that lends itself to a dimensionally stable scaffold in the presence of osteoconductive fillers (such as hydroxyapatite), and the final material can also be rendered porous. Rather than relying upon the inclusion of salts [7], effervescent fillers that generate CO2 as the material cures were exploited to yield porosity. This PPF foam has been previously characterized both in vitro [8] and in vivo [9]. The feasibility of this approach was determined not only by the capacity of the nanoHA to demonstrate enhanced cell attachment, migration, and proliferation in vivo, but also by its mechanical ability to accelerate the reconstructive process of bone repair. The experimental in vitro and in vivo data presented herein were used to establish outcomes for feasibility based on the histologic sequences of healing of bone defects and supported establishing material functionality. II CLINICAL MANAGEMENT OF MANDIBULAR AND MAXILLOFACIAL DEFECTS Bone grafting procedures have become almost an integral part of implant reconstruction. In many instances, a potential implant site in the upper or lower jaw does not offer enough bone volume or quantity to accommodate a dental implant. This was usually a result of bone resorption that had taken place because one or more teeth (if not all) were lost. Bone grafting procedures generally try to reestablish bone dimension that was lost due to resorption.

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In the past, lack of bone posed a considerable problem, and sometimes implant placement was impossible. Newer techniques, however, offer the opportunity to place implants of proper length and width and to better restore functionality and esthetic appearance. Grafting material can be grouped into five different categories: (1) autograft or autogenous bone graft, (2) allograft or allogenic bone graft, (3) xenograft or xenogenic bone graft, (4) alloplast or alloplastic bone graft, and (5) growth factors. The autograft is considered the gold standard of grafting materials. For most grafting purposes confined to oral implantology, another part of the jaw (i.e., chin or back portions of jaw) is generally chosen as a donor site. This allows surgery to stay inside the mouth, thus avoiding any extraoral wounds and scarring. Sometimes, however, when there is not enough bone volume available intraorally, iliac crest bone is chosen as a secondary donor site. Allograft material is generally taken from cadaver bone due to its unlimited availability. This bone however must be treated in order to render it neutral to immune reactions and to avoid cross-contamination of host diseases. These treatments may include irradiation, freeze-drying, acid washing, and other chemical treatments. Xenograft, on the other hand, is often of bovine origin. Tissue banks usually choose these graft materials because it is possible to extract larger amounts of bone with a specific microstructure, an important factor for bone regeneration, as compared to bone from human origin. The alloplast usually includes any synthetically derived graft material not derived from animal or human origin. In oral implantology, this usually includes hydroxyapatite or any formulation thereof. III COMMON MANDIBULAR AND MAXILLOFACIAL BONE GRAFTING PROCEDURES A Sinus Augmentations One of the most frequently applied grafting procedures is the sinus augmentation. This procedure is restricted only to the upper jaw. With aging, the pneumatization of the para-nasal sinuses often occurs. Once teeth have been lost in that particular area it is difficult, if not impossible, to place endosseous implants, as shown in Fig. 1. For this particular problem, grafting methods have been developed that raise the bottom of the sinus by grafting bone underneath and thus creating enough space for one or more dental implants. This procedure has been performed successfully for more than two decades and is considered both acceptable and predictable. The grafting material used can be any one of the five categories previously mentioned. Again, autogenous bone generally gives the best and fastest results. However a considerable volume of bone (5 to 10 cc per side) is required to perform a typical sinus augmentation, usually more than can be harvested from an intraoral donor site. Therefore, use of an allograft, alloplast, xenograft, or a combination of the three is sometimes necessary. An autograft takes approximately 4 to 6 months to mature in the sinus compared to an allograft, alloplast, or xenograft, which may take 9 months or more. Sinus augmentations and implant placement can sometimes be performed as a single procedure, if enough bone between the upper jaw ridge and the bottom of the sinus is available to stabilize the implant. If insufficient bone is available, the sinus augmentation will have to be performed as a two-part procedure allowing the graft to mature for several months (depending on the graft material used) before the implants can be placed. For this type of case, a mechanically sound biodegradable material would be extremely useful in augmenting autologous bone grafting.

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Figure 1 Top: the para-nasal sinuses of a young patient are outlined. Here the sinuses are not very large. Notice the distance between the bottom of the sinus and the top of the ridge. Bottom: the image shows a toothless lower jaw with two rather large sinuses on each side of the nose (middle structure). Notice that there is virtually no space remaining between the top of the ridge and the bottom of the sinus.

B Onlay Grafts This type of grafting procedure is designed to re-establish bone which has been lost due to resorption brought on by previous tooth loss in that area. Commonly, several pieces of autogenous bone (usually taken from the chin or the very back of the lower jaw) is attached to the site of the bone deficiency. The area is then closed, and after a maturing period this piece of bone is incorporated into the host bed and solidly fused such that implants may be securely placed. Larger areas of resorption will need to be augmented with more pieces of autogenous bone. For those cases, the patient’s bone is harvested from the iliac crest or tibia. With se-

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vere bone resorption, other implant modalities can sometimes be chosen to circumvent this rather aggressive surgical approach. C Ridge Expansion This is a technique used to restore lost bone dimension when the jaw ridge gets too thin to place conventional root form implants. In this procedure, the bony ridge of the jaw is literally expanded by mechanical means. A series of expanders (in cross-section round or Dshaped metal rods of successively increasing diameter) are being forced into the chosen implant site. This is accomplished by tapping these expanders into the ridge with a surgical mallet. This will compress the inner spongy part of the bone and, if done properly, will bulge out the outer cortex. At this point, an appropriate implant can either be established immediately into the created socket, or a bone graft can be placed and allowed to mature for a few months before placing the implant. From a biomaterials perspective, the challenge has been to develop a material that encourages the reconstructive process by supporting a high density of ingrowing cells within the scaffold, diffusion of factors between the transplanted cells and the surrounding tissue following implantation, vascularization, and cosmetic recovery [10]. The PPF-based scaffold may meet this challenge by presenting sufficient hydrophilicity for cell attachment and proliferation, offering demonstrable porosity for cellular migration, contributing a richness of surface area for neovascularization, and providing sufficient dimensional stability for support of the reconstructive process.

IV TISSUE REGENERATION USING POLYMER SCAFFOLDS Bone is considered by some to be the prototypic model for tissue engineering [1]. Poly(propylen fumarate) foam has been investigated as a porous scaffold that will function as a cell and tissue carrier emphasizing material superstructuring. Therefore, this study was aimed at engineering a biodegradable superstructure that would provide optimal spatial and nutritional conditions for cell maintenance by the arrangement of structural elements (e.g., pores) that vary the order of cell-to-cell contact [11]. Predictable three-dimensional cell– polymer matrices for tissue engineering are needed so that osteoblast cells can maintain their phenotypic properties and form a mineralized matrix while seeded on the polymer surface [2,12]. Optimization of a predictable and dimensionally stable scaffold maintaining the three-dimensional superstructure of the varying reconstructive sites in combination with the hydroxyapatite nanoparticle technology has been at the center of this developmental approach. Bone cells can be readily cultured from bone tissues [7,13,14]. These cells have been grown in synthetic polymeric as well as in natural matrixes and are, therefore, recognized as an extremely bioactive material [15]. It is logical to take bone cells from the future recipient of the bone repair material and to expand them in vitro for engineering of repair tissues. Breitbart et al. [16] have already demonstrated the feasibility of using periosteal cells for tissue engineered bone repair of calvarial defects. Ishaug et al. [17,18] have shown that osteoblasts can grow and migrate throughout polymeric scaffolds in vitro. Ishaug-Riley et al. [7] further demonstrated that osteoblast function on synthetic materials was equal to that on nondegradable orthopedic materials. Seeded resorbable scaffolds were replaced by new bone when implanted into bony sites.

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V HYDROXYAPATITE COMPOSITIONS AND NANOPARTICLE TECHNOLOGY Hydroxyapatite, Ca10(PO4)6(OH)2, is an attractive and widely utilized bioceramic material for orthopedic and dental implants that closely resembles native tooth and bone crystal structure. Though hydroxyapatite is a common bioceramic, applications for its use have been limited by its processability and by its architectural design conceptualization. Conventional processing produces hydroxyapatite that lacks compositional purity and homogeneity. Because it has been difficult to sinter, dense hydroxyapatite structures for dental implants and low-wear orthopedic applications typically have been obtained by high-temperature and/or high-pressure sintering with glassy sintering aids that frequently induce decomposition to undesirable phases with poor mechanical stability and poor chemical resistance to physiological conditions. Thus, conventionally formed hydroxyapatite necessitates expensive processing and compromises structural integrity due to the presence of secondary phases. Existing methods require high forming and machining costs to obtain products with complex shapes. Furthermore, typical conventional hydroxyapatite decomposes above 1200°C. This results in a material with poor mechanical stability and poor chemical resistance. Jarcho et al. [19] have described a process for forming dense polycrystalline hydroxyapatite that is “substantially stronger than other hydroxyapatite materials” and that elicits “an excellent biological response when implanted in bone.” A precipitation method was used and material of average grain size of 150–700 nm had been recovered. However, they have reported low volume fraction of pores and considerable grain growth during sintering, even at firing temperatures of 1000°C [19]. Jarcho et al. [19] has achieved 99% density in some cases, but used a technique that can be impractical for forming desired shapes. Akao et al. [20] have reported the compressive flexural and dynamic torsional strengths of polycrystalline hydroxyapatite sintered at 1300°C for 3 hours and have compared the mechanical properties of the product with those of cortical bone, dentine, and enamel. They reported that compressive strength of the sintered hydroxyapatite was approximately three to six times as strong as that of cortical bone. There is much room for improvement in the use of hydroxyapatite as implants as reported by Hench et al., [22] “because (hydroxyapatite) implants have low reliability under tensile load, such calcium phosphate bioceramics can only be used as powders, or as small, unloaded implants such is in the middle ear, dental implants with reinforcing metal posts, coatings on metal implants, low-loaded porous implants where bone growth acts as a reinforcing phase, and as the bioactive phase in a composite.” Hench et al. have also reported that hydroxyapatite had been used as a coating on porous metal surfaces for fixation of orthopedic prostheses. In particular, hydroxyapatite powder incorporated into the pores of porous, coated-metal implants would significantly affect the rate and vitality of bone ingrowth into the pores. Plasma spray coating of implants generally has been the preferred technique for applying a thin even coating to metal implants. Hench et al. have reported, however, that long-term animal studies and clinical trials of load-bearing dental and orthopedic prostheses have suggested that the hydroxyapatite coatings may degrade or delaminate [22]. Thus, the creation of new forms of hydroxyapatite with improved mechanical properties would have significant use. Recent attention has been focused on nanocrystalline or nanocomposite materials for mechanical, optical, and catalytic applications. By designing materials from the cluster level, crystallite building blocks of less than 10 nm are possible from which unique size-

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dependent properties such as quantum confinement effect and superparamagnetism have been obtained. Various nanocrystalline ceramics for structural applications were rigorously investigated during the 1990s. Siegel et al. [23] have discussed nanophase metals and ceramics noting that although many methods exist for the synthesis of nanostructured materials—including chemical or physical vapor deposition, gas condensation, chemical precipitation, aerosol reactions, and biological templating—synthesis and processing methods for creating tailored nanostructures are still sorely needed, specifically techniques that would allow careful control of surface and interface chemistry and that would lead to adherent surface coatings or well-consolidated bulk materials. In the case of normally soft metals, it has been found that decreasing grain sizes of the metal below a critical length scale (less than about 50 nm) for the sources of dislocations in the metal increases the metal’s strength. Clusters of metals, intermetallic compounds, and ceramics have been consolidated to form ultrafine-grained polycrystals that have mechanical properties remarkably different and improved relative to their conventional coarse-grained counterpart. Nanophase copper and palladium, assembled from clusters with diameters in the range of 5–7 nm, have been noted for having hardness and yield strength up to 500% greater than conventionally produced metal. It has also been noted that ceramics and conventionally brittle intermetallics can be rendered ductile when synthesized from clusters with sizes below 15 nm, the ductility resulting from the increased ease with which the ultrafine grains can slide by one another in grain-boundary sliding. However, synthesis of nanocrystalline or nanocomposite materials has been difficult. Significant effort has been put into such synthesis, and it is likely that most attempts for producing particle sizes on the nanometer scale do not result in a high yield due to agglomeration. A delicate balance of synthetic parameters typically must be elucidated in conjunction with a particular set of materials. Nanocomposites including intra- and intergranular nanocomposites and nanonanocomposites demonstrated improvement of mechanical properties, machinability, and superplasticity. Although hydroxyapatite has been widely used, and a hydroxyapatite formulation having mechanical and morphological properties advantageous for prostheses would be very useful, production techniques to date have not always produced reliable structural implants. VI PREVIOUS STUDIES WITH PPF FOAM SCAFFOLDS A dimensionally stable porous scaffold was investigated based on the crosslinking of the unsaturated PPF polymer with a vinyl monomer, vinyl pyrrolidone (VP), in the presence of both effervescent sodium bicarbonate and citric acid fillers and the osteoconductive filler hydroxyapatite (HA). Upon mixing, the polymer cures via crosslinking of the PPF by the monomer and concomitant CO2 generation resulting in a porous scaffold degradable by hydrolysis. The use of hydroxyapatite as part of the filler supports the osteoconductivity of the scaffold [24], whereas the CO2 generated pores provide porous regions for attachment and proliferation of cells not only in situ, but ex situ; the hydrophilicity of the polymeric support encourages cellular migration. This PPF scaffold has been evaluated in vitro for morphological, mechanical, and surface properties and in vivo using a rat tibial defect model [25]. Scaffolds, designed for controlled superstructural characteristics, have been reported to be hydrophilic, mechanically comparable to trabecular bone (6.4 MPa compressive strength and 130 MPa compressive modulus), dimensionally stable (Figs. 2 and 3), and porous [9]. Histologic and histomorphologic examination of the implant region of rats has suggested that the porosity of

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Figure 2 Scanning electron micrograph of PPF foam at t  0.

Figure 3 Scanning electron micrograph of PPF foam after 24 days in vitro (PBS, pH 7.23, 37°).

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the scaffold supported bony ingrowth, and the stability of the scaffold preserved the dimensional integrity of the defect site. A Experimental Objectives and Rationale This study attempted to determine the feasibility of implanting dimensionally stable, bioresorbable foams fashioned from a crosslinked biopolymer, poly(propylene glycol-co-fumaric acid) that has been thought to be suitable for through-and-through cell penetration by inclusion of nanoparticulate hydroxyapatite to enhance bone reconstruction. The development of a designable reconstructive material with osteoconductive properties supporting the actual three-dimensional structure of clinically applicable tissue equivalents focused on implantation into a small animal (rat) model. In analogy, the proposed clinical approach would eventually involve (1) the procurement of autologous bone graft from the recipient, (2) the augmentation of a nanoparticle-loaded biopolymeric foam with these autologous cells, and (3) the transplantation of these bioprocessed scaffolds into a recipient. Ideally, this new approach to biodegradable bone graft replacement materials would additionally promote enhanced bony ingrowth by utilizing nanoparticle hydroxyapatite. Therefore, the bioactivity of nanohydroxyapatite has been evaluated when employed in a bioresorbable bone graft substitute and implanted into a bony defect model. For this study, results of comparative histologic and histomorphometric analysis of bone ingrowth after implantation of nano- and microhydroxyapatite using poly(propylene glycol-co-fumaric acid) scaffold is reported and results presented indicate enhanced healing when nanohydroxyapatite composites were used. B Experimental Design 1 Formulations The general formulation used for the study is shown in Table 1. The PPF (Mw ~5,000 by GPC) was synthesized from equimolar fumaric acid and propylene glycol in the presence of p-toluene sulfonic acid [39,40]. 1-Vinyl-2-pyrollidinone (VP) was purchased from Aldrich (USA) and used after vacuum distillation. Benzoyl peroxide (BP), hydroquinone (HQ), and N-N-dimethyl-p-toluidine (DMPT) were purchased from Aldrich and used as received. The PPF was crosslinked with liquid VP, using BP (2.1%w/w) as the initiator. The accelerator, DMPT, at a concentration of 0.03%w/w, gave a working time of approximately 90 s. The formulations were injected into Teflon-lined multiple-well cylindrical molds, each well measuring 6 mm in diameter by 12 mm in length, and allowed to cure at 37°C. Table 1 General Composition of the Two-Part Formulation Part I PPF HA SB CA BP Total

Wt. (mg)

Wt. (%)

Part II

Wt. (mg)

Wt. (%)

1179.5 341.5 51.3 43.4 52.5 1668.2

(47.2) (13.7) (2.1) (1.7) (2.1) (66.8)

VP DMPT H2O Total

380.0 0.65 450.0 845.65

(15.2) (0.03) (18.0) (33.2)

Note: PPF, poly(propylene glycol-co-fumaric acid); HA, hydroxyapatite; BP, benzoyl peroxide; SB, sodium bicarbonate; CA, citric acid; VP, vinyl pyrrolidinone; DMPT, dimethyl p-toluidine; H2O, distilled water.

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The PPF foam with pore sizes of 100–300 m appeared to be desirable for bone cell ingrowth. Bondre et al. [9] have demonstrated that the design of the porosity of the foam was attainable by control of the sodium bicarbonate (SB)/citric acid (CA) content and by the size of the SB/CA particles used in the effervescent filler. The reaction of CA/SB with water produces carbon dioxide, the blowing agent responsible for foam formation and expansion. The stoichiometry requires a 1:3 mole ratio of CA/SB with a corresponding weight ratio of 1.00:1.31. The moles of CO 2 which can be generated per gram of material depend on the loading of CA/SB in the foaming cement. A 0.15% CA/SB loading would produce a 25% expansion at 37°C and 1 atm based on this stoichiometry. In addition to the blowing agent, the PPF formulation was crosslinked using vinyl pyrollidinone in the presence of an osteoconductive HA filler employing techniques described previously by Lewandrowski et al. [9]. The two types of HA investigated were (group 1) sintered, spherical micro-HA (median particle size  26 m, commercially available from CAM Implants, The Netherlands) and (group 2) sintered, spherical nano-HA (median particle size  40 nm, as produced and characterized by Ying et al. [26–32]. All HA preparations used in this study have been examined using x-ray diffraction (XRD) to characterize the crystalline purity and size, photoacoustic Fourier transform infrared (PA-FT-IR) spectroscopy to substantiate the molecular structure, and transmission electron microscopy (TEM) to determine the particle size and morphology. C In Vivo Animal Studies and Group Design Three groups were tested in the rat tibial metaphysis implantation model according to Gerhart et al. [25]: the nano- and micro-HA available (groups 1 and 2) and the unfilled control (group 3). [33]. Adult male Sprague Dawley rats weighing approximately 200 g were used as the animal model (Zivic Miller, Zelienople, PA). Animals were anesthetized using an intramuscular injection of ketamine HCl (100 mg/kg) and xylazine (5 mg/kg). The rats were also given an intramuscular prophylactic dose of penicillin G (25,000 U/kg), and the surgical site was shaved and prepared with a solution of Betadine (povidone-iodine) and alcohol (Dura-Prep, 3M Health Care, St. Paul, MD). A 1.5-cm longitudinal incision was made in the anterior left hind leg, and the tibial metaphysis was exposed. A 3-mm hole was made in the anteromedial tibial metaphysis of rats. The formulations, mixed immediately prior to surgery to a consistency similar to a paste or putty, was implanted into the prepared tibial defect site using a spatula. The PPF-based grout cured in situ and, after the implantation of the bone grout, the soft tissues and skin were closed in layers with running absorbable sutures. One batch of each formulation was implanted in each of eight animals. Thus for the three groups, a total of 24 animals were used. All the animals were sacrificed 3 weeks postoperatively. D Methods of Evaluation Evaluation was performed on high-resolution radiographs taken immediately postoperatively and at 3-week intervals until sacrifice using a specimen x-ray unit (Microfocus 50E6310F/G, Xerox, Rochester, NY). Radiographs were taken with minimal exposure (32 kvp, 2 s), and mammography film (Cronex Microvision, Dupont Medical Products, Wilmington, DE), cassettes (MR Detail, AGFA Richfield Park, NJ), and screens (Mammoray AGFA) were employed. Following sacrifice, 10-mm segments of the tibial bone including the section that was implanted with a bone graft substitute were harvested. The specimens

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were processed for histologic analysis by fixation in 10% buffered formalin. Specimens, which included residual bone graft material, were decalcified in EDTA and paraffin embedded. Longitudinal sections (5 m thick) of the total specimen were then cut and stained with hematoxyline and eosin. In addition, slides were stained according to the von Kossa method to isolate calcium crystals. Slides were examined for resorptive activity and new bone formation at the implantation site as well as for inflammatory responses. Histomorphometric evaluation of new bone formation around the different types of grafts was done by acquiring images of serial longitudinal hematoxyline and eosin stained sections of the specimen using a CCD video camera system (TM-745, PULNiX, Sunnyvale, CA) mounted on a Zeiss microscope. Images were digitized and analyzed using Image Pro Plus software. For each specimen, the area of newly formed bone surrounding the implant and within the implant was measured. This measurement was standardized against the total area occupied by the implant in the same section. A minimum of five sections obtained from different levels of the specimen were included for this analysis. The spacing between sections of adjacent levels was typically 300 m. This allowed for an approximate absolute volume of the newly formed bone, given as an average percentage rate (mean  standard deviation) for each bone specimen. To compare the extent of new bone formation around the graft at its metaphyseal implantation site between the experimental groups and the control group, the remodeling index was determined. This was defined as the volume ratio of newly formed bone and the volume of the whole cortical allograft based on eight animals per study group. These were also given as average percentage rates. Differences in the remodeling indices were analyzed for statistical significance by employing an ANOVA test. A p-level of 0.05 was considered statistically significant. E Preliminary Results 1 Control Group In the control group (group 3), new bone formation was not observed in the metaphyseal defect 3 weeks postoperatively. There was some periosteal bone formation at the cortical drill hole site, but the remainder of the defect in the tibial metaphysis was filled primarily with bone marrow and fatty tissue. In the microhydroxyapatite group (group 2), the implant remained structurally stable and did not disintegrate. There was no histologic evidence for implant dissolution or active cellular resorption from the recipient site. Although there was some moderate infiltration with PMNs, this appeared to be consistent with postoperative inflammatory changes. In addition, new bone formation was observed, which at 3 weeks postoperatively appeared to be tightly packed around the implant without excessive fibrous or inflammatory tissue. There was osteoclastic and osteoblastic activity at the surface of the implant suggesting that the bone surrounding it was undergoing active remodeling. Although the bone surrounding the implant bone graft underwent active remodeling, the implant remained structurally intact (Fig. 4). Microhydroxyapatite crystals were easily demonstrated with the von Kossa stain. New bone formation and large round cells whose morphologic appearance was consistent with osteoblasts were noted in close proximity to the micro-HA crystal (Fig. 5). In the nanohydroxyapatite group (group 1), the implant surface appears to have stimulated a more vigorous inflammatory response with infiltration by PMNs and macrophages. In addition, an increase in new bone formation was observed around implants containing nanohydroxyapatite. As in the microhydroxyapatite group, the implant remained structurally stable and did not disintegrate. There was no histologic evidence for implant

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Figure 4 Photomicrograph of longitudinal sections (H&E) of a rat tibia procured at 3 weeks postoperative. The drill hole defect was filled with a PPF scaffold containing micron-sized hydroxyapatite showing some reactive bone formation around the PPF-based implant (10). Note the width of the newly formed bone margin surrounding the implant is on the order of 50 to 150 m.

Figure 5 Photomicrograph of longitudinal sections (H&E) of a rat tibia procured at 3 weeks postoperative. The drill hole defect was filled with a PPF scaffold containing micron-sized hydroxyapatite. This high power photograph (20) shows micron-sized HA within newly formed bone. Multiple osteoblast depositing osteoid onto the implant are also shown.

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Table 2 Histomorphometric Analysis of New Bone Formation Groups

Remodeling index (%)

Empty defect Micro-HA PPF implant Nano-HA PPF implant

12.3  6.9 34.3  10.6 48.3  14.9

dissolution or active cellular resorption from the recipient site. In contrast to the microhydroxyapatite group, HA crystals were not stainable with the von Kossa technique in the nanohydroxyapatite group. Similarly, as in the micro-HA group, large cells with a round nucleus positioned toward the interface with the implant were present. Osteoids appeared to be secreted onto the implant material. Histomorphometry showed that the amount of new bone that formed around the different types of grafts used in this study was significantly higher in the nanohydroxyapatite group than in the control group (no implant; p  0.002) and in the micro-HA group (p  0.025). Although both formulations were equally osteoconductive, as measured by the implant area covered by newly formed woven bone, a wider margin of newly formed bone was noted around nanohydroxyapatite implants. Also, more new bone was found within these types of implants. As a result, the remodeling index was higher for the nanohydroxyapatite group when compared to the microhydroxyapatite group (see Table 2). VI CONCLUSIONS Nanohydroxyapatite could eliminate some of the disadvantages associated with the use of conventional hydroxyapatite. It has been produced in a purer, more homogeneous form. Copolymer composites have proven to have better mechanical properties when nanoapatite particles were used. Liu et al. [34] hydrothermally synthesized acicular nanoapatite (Nap) that was used as filler to make composites with a polyethylene glycol/poly(butylene terephthalate) (PEG/PBT) block copolymer (Polyactive 70:30). The Nap had a particle diameter of 9–25 nm and a length of 80–200 nm. The mechanical properties and the physicochemical characteristics of the composites, such as Young’s modulus, swelling degree in water, and the calcification behavior, were determined. Nanoapatite had a strong ability to promote the calcification of composites when incorporated into Polyactive 70:30. In the dry state, Nap had a prominent stiffening effect for Polyactive 70:30 [34]. If replicable in an aqueous environment, such reinforcement of the polymer by Nap would have immediate applicability for orthopedic implant and repair materials. To date, little is known about the biological effects of nanohydroxyapatite on the healing process of bone defects. Investigators have analyzed the effect of nanohydroxyapatite on osteogenic cells in vitro and have noted an increase bioactivity. Du et al. [35] produced a porous nano-HA/collagen composite (nHAC) in sheet form, which was convolved to be a three-dimensional scaffold. Using organ culture techniques, they developed three-dimensional osteogenic cells/nHAC constructs in vitro. The porous nHAC scaffold was noted to provide a microenvironment resembling that seen in vivo, and cells within the composite eventually acquired a three-dimensional polygonal shape. In another study, Du et al. [36] investigated the tissue response to a nanohydroxyapatite/collagen composite implanted in a

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marrow cavity. Histologic and scanning electron microscopic evaluation demonstrated the material to be bioactive as well as biodegradable. At the interface of the implant and marrow tissue, solution-mediated dissolution and giant cell–mediated resorption led to the degradation of the composite. Bone formation at the periphery of the implant was also evident. The process of implant degradation and bone substitution appeared to be reminiscent of bone remodeling. The ultimate objective of this animal study in rats was to establish the utility of nanohydroxyapatite in an osteoconductive grout while demonstrating biocompatibility in the absence of foreign body reactions and evaluating the effect on the bone healing/repair process. The PPF-based resorbable bone graft substitute presented here was expected to be osteoconductive because the hydroxyapatite filler has been successfully employed in similar model evaluations. In this initial biocompatibility study in rats, the histomorphology of the osteointegration process of a PPF-based bioresorbable bone graft material augmented either with conventional, micro-HA, or nano-HA was investigated. Two types of HA particle sizes were studied: (1) nano-HA (median particle size  40 nm) and (2) micro-HA (median particle size  26 m). These formulations were tested in an animal model developed and utilized by Gerhart et al. [25]. This screening model allowed facile evaluation of the grouting process and easy visualization of tissue bonding, in addition to allowing comparative histologic assessments of degradation and bone ingrowth. New bone formation in the negative control group (group 3), was absent, i.e., consisting of unfilled defects. At 3 weeks, more reactive new bone formation was observed in the nano-HA group than in the micro-HA group without complete closure of the defect for either of the two test groups. These histologic observations were supported by histomorphometric measurements of new bone formation that demonstrated significant increases when PPF was used in combination with nano-HA (Table 1). For the duration of the study (3 weeks), no evidence of implant failure or disintegration was observed. In addition, it was clearly evident that these PPF-based bone graft substitutes were extremely osteoconductive, showing both ingrowth of newly formed woven bone with concurrent neovascularization. This is consistent with previously reported investigations, which suggested osteoconductive properties of biopolymers could be improved by the addition of hydroxyapatite [37,38]. The accompanying inflammatory responses appeared more pronounced in the nano-HA group than in the micro-HA group. This result was consistent with reports by Du et al. [36], who noted similar histologic observations. It remains unclear, however, what the basis of this sustained inflammatory responses was and whether or not it was of any significance for the bone healing/repair process. Based on these observations, this study has suggested that the osteoconductive properties of a PPF-based bone graft material containing 13.7% HA could be further improved by utilization of nanohydroxyapatite. Rapid bony ingrowth and healing could be facilitated by accelerated bone formation around and within a biodegradable scaffold. Further investigation of nanohydroxyapatite should focus on designing bone graft materials that are structurally superior to established bone graft substitute materials and capable of stimulating in vivo ingrowth whose temporal sequence is synchronized with the sequence of histologic events of the bone healing. ACKNOWLEDGMENTS The authors wish to thank Dr. Joseph Alroy, Associate Professor in Pathology, Tufts University Schools of Medicine and Veterinary Medicine for his assistance in the histologic

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analysis of this study. This work was supported in part by NIH/NIDCR Grant No. 5 R44 AR 44317 (to Joseph D. Gresser) and NIH/NIAMS Grant AR 45062 (to Kai-Uwe Lewandrowski).

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19 Segmental Bone/Joint Replacement Using Guided Tissue Regeneration Edmund Y. Chao, Nozomu Inoue, and Frank J. Frassica Johns Hopkins University, Baltimore, Maryland Frank H. Sim Mayo Clinic and Mayo Foundation, Rochester, Minnesota

I INTRODUCTION Due to recent advances in the staging of musculoskeletal sarcomas, improvements in diagnostic imaging, and the use of adjuvant and neoadjuvant therapies, orthopedic oncologists can now offer limb salvage surgery to up to 90% of patients as an alternative to amputation [36,39,46,50,57]. Studies have shown a significant improvement in the survival and local control rates of patients with malignant bone tumors due to adjuvant chemotherapy and improved surgical techniques [3,101]. The survival rate for patients with limb-sparing procedures was found to be similar to those treated by amputation [83,93,94]. Hence, in carefully selected cases, limb salvage does not appear to adversely affect the long-term prognosis, even in the presence of high-grade malignancy. When local resection and limb salvage are feasible, a variety of surgical techniques have been used, including letting the limb segment flail, reconstructing the defect with an autogenous graft [37,95,105] or an allograft [1,2,6,8, 25,38,64–66,69,70,76], and using a custom-designed or a modular prosthetic device [7,9, 19,22,34,35,62,63,88,91]. Each method has its advantages and disadvantages. Prosthetic replacement was found to be a promising alternative to other limb reconstructive methods with many potential advantages. However, implant loosening, component fracture, and the potential biological reaction to metallic and polyethylene particles remain the long-term problems [15,16]. The use of custom-designed segmental bone/joint prostheses to reconstruct skeletal defects and restore joint function has become the dominant reconstructive method in or355

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thopedic oncology. However, the patient’s functional results and the prosthetic device performance need further improvement, as the current prosthetic failure rate at 5–10 years follow-up has been reported at 35 to 40% [17,77]. Porous coating technology was adopted early in our stem design for cementless intramedullary fixation, but the results were disappointing due to loosening, bone resorption, and stem fracture. Additional failures were attributed to loosening, dislocation, and infection [82,89]. Other problems were related to implant availability at the time of surgery, improper size, and high cost. Improvements in the off-the-shelf modular prosthetic design and the development of more durable stem fixation to bone were achieved and have significantly improved medium-term results. Modular systems containing fewer interchangeable parts and with the porous coating confined to the body of the implant were later developed to alleviate these problems (Fig. 1). An innovative concept of utilizing cemented solid intramedullary stems to provide the initial and

B

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Figure 1 The modular segmental bone/joint replacement prosthesis systems developed by the senior author and several orthopedic oncologic surgeons. (A) The schematic diagram of the modular oncology system made of Co-Cr alloy and coated with Co-Cr beads made by Howmedica since 1986. (B) The stem component with porous coating in the shoulder region of the body segment. (C) The modular segmental prosthesis with Ti fibermetal coating for proximal humeral replacement made by Zimmer since 1979. The system contains all major anatomical regions of replacement, but the implant is available only through custom service.

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Figure 2 The composite fixation method defined as extracortical bone bridging and ingrowth (EBBI) following the principles of guided tissue regeneration.

short-term fixation, while bone formation and ingrowth over the porous prosthetic segment increased the fixation strength through a biological re-enforcement was incorporated. This “composite” fixation concept was termed extracortical bone bridging and ingrowth (EBBI) fixation (Fig. 2). To further refine this new implant design and fixation concept, several additional issues were addressed: (1) Is the conical coupling joint in the modular system strong enough to sustain repeated loading without fracture, loosening, or wear particle generation? (2) Could effective autogenous graft substitutes be used to enhance the quality of EBBI? (3) Would the formation of EBBI be affected by adjuvant cancer therapies? (4) Would the stress transfer across the bone bridging site remain effective for EBBI maintenance without resorption? (5) Could the application of stem cell–rich bone marrow or rhBMP provide additional enhancement to the formation and maintenance of EBBI? Animal experiments, bench tests, and analytical studies were performed to address these questions in a systematic manner. These study results plus the long-term patient follow-up data have led to a firm foundation of this stem fixation concept for segmental bone and joint replacement prosthesis application. Reattachment of resected soft tissues onto the metallic prosthesis surface has been a challenge in musculoskeletal reconstructive surgery. Fixation of host tissue to the segmental prosthesis using staples, washers, and metallic loops has not been effective in the past. For this reason, an alternative method using a special porous tendon anchoring device (the enhanced tendon anchor, ETA, device) was first developed in 1986. The ETA device can be attached to various segmental replacement prostheses in proximal femur, proximal humerus, and proximal tibia [60]. Attaching the ETA device to the segmental prosthesis may allow direct tendon attachment to the original anatomical site, resulting in better muscular function and joint stability. The current development follows the working hypothesis that for secured and functional results of tendon attachment to metallic implant, normal morphology of direct tendon attachment to bone must be duplicated in accordance with the basic principles of tissue engineering. Strong support of this bold attempted was received from a long series of in vitro and in vivo studies and the recent development of connective tissue regeneration growth factors.

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These exciting concepts were supported by the principles of guided tissue regeneration introduced by the dental and oral surgery field, which predated the introduction of tissue engineering by at least 6 years [73]. The biological principles of guided tissue regeneration were established as a method to repair and replace lost periodontal tissue. Later, the same method was used to heal long bone defects [28,29]. The basic idea behind this method was that the use of biocompatible synthetic materials to form a membrane to exclude unwanted tissue in a defect space, and the normal tissue would gradually form to fill up the defect. The materials used for the membrane were Gore-Tex, resorpable PLA, e-PTFE, decalcified freeze-dried allograft, titanium, etc. [27,41,49]. In most periodontal work, the regenerated tissue served the purpose of space filling with little or no functional load-bearing requirement. In orthopedic surgery, however, stability of the defect construct must be maintained using prosthetic scaffolds while osseous tissue could be recruited or guided to provide secured and permanent biological reconstruction. This was the underlying working principle behind EBBI and was later adopted for the direct tendon attachment to metallic implants using the ETA augmented by the grafting material from the host. As the autogenous grafting material is always limited in supply and the recipient sites are usually affected by adjuvant therapy in tumor patients or the scar tissue with poor vascular condition in trauma or total joint replacement revision cases, more reliable and effective enhancement factors must be added to assure the success of these composite implant and tendon fixation concepts [23]. Such need of biological factors appears to be an ideal and fitting application of the tissue engineering principle where externally derived or engineered growth factors could be applied to regenerate connective tissue. The term tissue engineering was first coined by Y. C. Fung in a National Science Foundation bioengineering workshop in 1988 [68]. Tissue engineering was formally defined as “the application of the principles of engineering and the life sciences towards the development of biological substitutes that restore, maintain, or improve tissue function” [61,79]. This is a rather broad definition which would encompass many fields in orthopedic surgery including fracture healing, bone lengthening, osteotomy, bone and joint replacement, as well as the concept of guided tissue regeneration, all of which rely on patient’s own biologic potential. The more attractive alternative would be the use of derived or engineered proteins and growth factors which could be delivered in situ and released at an appropriate rate to enhance the formation and maintenance of specifically targeted connective to fulfill their underlying functional requirements. These biological factors combined with sound implant design and effective surgical technique will be essential to realize the goals of EBBI in prosthetic implant fixation and for the direct tendon attachment to metallic implant to improve patient’s function and reduce complications such as articular surface wear or joint dislocation. II BIOMECHANICAL AND BIOLOGICAL JUSTIFICATION OF EBBI FIXING The conceptualization of the extracortical bone bridging and ingrowth idea was entirely a serendipitous process after lengthy observation of the extensive usage of cemented custom segmental bone/joint replacement prosthesis from 1972 and the complications experienced in the early use of porous-coated stem implant during the period of 1975–1979 [10]. This concept preceded the guided tissue regeneration method by several years. The use of autogenous bone grafting to induce the formation of extracortical bone bridging across the implant shoulder and bone host followed the well-established osteogenetic principle using the

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onlay bone graft first introduced by Phemister in 1947 [78]. The clinical trial of the EBBI implant fixation method with custom-made prosthesis was conducted simultaneously to the basic science investigations. Under successful bone bridging and ingrowth, the stresses in the cement and stem were significantly reduced (Fig. 3). Such stress bypassing effect was also effective even where the bony bridging around the implant shoulder region was incomplete. Such biomechanical benefit was still significant when the bridging tissue was primarily fibrous. Although the bone stress was not significantly increased following successful EBBI, this could be caused by the assumption of rigid bonding interface between bone and cement and between stem and cement. In reality, there will be micromovement occurring at the interfaces, which would allow more stress transmitting through host bone cortex since there was no clinical finding of bone resorption in any case of successful EBBI regardless of the length of follow-up.

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Figure 3 (A) Stem stress and (B) cement stress as affected by EBBI and the type of tissue formation involved. Three-dimensional finite element analysis technique was used with unit applied load in torsion, axial compression, and bending.

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Extracortical bone formation over the prosthesis has been achieved with autogenous iliac cortical and cancellous bone graft. This prosthetic fixation concept has been validated by animal experiments [71,74]. The short- and long-term clinical results showed significant improvement over the previous nonporous or porous-coated prostheses in regard to longterm stem fixation [18,47]. The main advantages for this composite fixation (cemented intramedullary stems and extracortical bone formation) are (1) cement provides secure initial fixation to enhance bone graft incorporation and allows early weight bearing and range of motion; (2) a solid stem increases implant fatigue strength; (3) fewer stem sizes facilitate the design of the modular system; and (4) the implant is easier to remove and revise, if necessary. Through finite element stress analysis and bench tests, the current taper joint design was optimized to provide sufficient locking strength to prevent uncoupling under different loading conditions [20,24,30,31]. In segmental bone and joint replacement, substantial bending moments are present during load bearing which can create significant tensile stresses in the critical regions of the taper lock [40]. In addition, the microgap opening of the taper joint interface could cause severe wear and allow body fluids to seep through, causing corrosion. The corrosion is increased when the modular components involve different types of metal [26,87]. If the adjacent shoulder region of the taper lock joint were porous coated and connective tissues formed a closed capsule, the stress levels within the joint components would be significantly reduced. In addition to load sharing by the EBBI, the “purse string” of soft tissue (Fig. 4) confines the metallic wear particles within the capsule and reduces the risk of “particle disease,” which can lead to bone resorption and stem loosening [21,102]. The experimental data and analytical models also showed that porous coating surrounding the Morse taper lock junction in the modular design of the segmental prosthesis would reduce bending stresses, minimizing fatigue failure of the joint component [11,12]. Soft tissue attachment along the prosthesis segment is expected to improve the control of the limb and thus the joint function.

Figure 4 The hypothetical stress bypass concept provided by the EBBI. The connective tissue also forms a tight capsule around the bone–implant junction, later described as the “purse string” effect to seal the stem/cement/bone interface from the wear particle invasion which may lead to osteolysis and aseptic loosening.

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III EFFECT OF PREOPERATIVE AND POSTOPERATIVE CHEMOTHERAPY ON EBBI There are many reasons for complications related to the use of prosthetic implants, including operative technique, wear debris, stress shielding, metal fatigue failure, and aseptic loosening of the methyl methacrylate at the bone–prosthetic interface. The concept of EBBI, using custom porous-coated prosthetic devices, has been developed to achieve a more durable reconstruction. Unfortunately, chemotherapy has an inhibitory effect on bone healing, bone ingrowth into prosthetic devices, and the incorporation of bone grafts [96, 103]. Clinical studies have shown an increased rate of complications in patients who have been on intensive multiagent chemotherapy regimens. Gait analysis, histomorphometrical evaluations, and scanning electron microscope evaluations of 24 canines have been performed to study the effects of pre- and postoperative cisplatin administration on the biological fixation of a porous-coated segmental prosthesis [107,108]. Eight dogs each were assigned to one of three treatment groups. Group 1 dogs served as sham-treated controls. Group 2 dogs were given cisplatin four times before surgical implantation of the segmental prosthesis. The first treatment was given 12 weeks before surgery, and subsequent treatments were given every 3 weeks. The last treatment was given 3 weeks before surgery. Group 3 dogs were given cisplatin four times after surgical implantation of the segmental prosthesis. The first treatment was given 1 week after surgery and subsequent treatments were given every 3 weeks. The last treatment was given 10 weeks after operation, 2 weeks before sacrifice. A new biomechanical torsion testing technique was developed to test only the EBBI tissue by removing the cement at the bone–stem interface using a special instrument (Fig. 5).

Figure 5 The special cement removal instrument to prepare the reconstructed femur to test the contribution of EBBI alone on prosthesis fixation.

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Figure 6 Single drug chemotherapy (cisplatin) effect on (A) extracortical bone formation and (B) torsional strength of the fixation.

The results showed that the chemotherapy did not have a statistically significant effect on postoperative weight bearing, and the mechanical parameters showed no significant differences between groups at 12 weeks (Fig. 6). Histologically, preoperative chemotherapy (group 2) did not alter bone formation, but the reconstructed limbs in group 3 (postoperative chemotherapy) had significantly less periprosthetic bone than group 1 (control) and group 2 (preoperative chemotherapy) specimens. Extensive osteoclastic and osteoblastic surfaces were observed. Cell activity was most notable around remaining particles of cortical bone graft. In the postoperative cisplatin-treated specimens, the cortical bone particles were larger than in other groups as a sign of less resorption and new bone formation. Bone ingrowth into the prosthesis did not differ significantly between groups. Doxorubicin, cisplatin, and ifosfamide have been used clinically for osteosarcoma patients. Though much debate exists regarding the best chemotherapy regimen, the combination has been shown to be effective [51,52]. These agents were used to develop a canine multiagent chemotherapy model for future studies (Fig. 7). The protocol consisted of three cycles of drug administration. Each cycle was planned to last 4 weeks, with adriamycin and cisplatin administered on the first day of the first week and ifosfamide administered on 4 days of the third week. For drug administration, an intravenous line was established in the

Figure 7 The multidrug chemotherapy canine model used to study its effect on EBBI and the segmental prosthesis fixation strength.

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cephalic vein. An intravenous infusion of 2 L of 0.9% NaCl was initiated at 20 mL/kg/h for prehydration. Next, adriamycin (ADR) was diluted in 250 mL of 0.9% NaCl and administered over a period of 20 min. After another 250 mL of 0.9% NaCl was also given at the same rate, cisplatin (CDP) diluted in 250 mL of 0.9% NaCl was also given over a 20-min period. At the conclusion of cisplatin administration, saline diuresis was continued using an additional 2 L of 0.9% NaCl at 20 mL/kg/h. On the third week of a cycle, ifosfamide (IFX) was given concomitantly with mesna at 1.5 times the concentration of ifosfamide for uroprotection. After the fourth and final day of ifosfamide administration for each cycle, mesna was given 8 and 16 h after ifosfamide at a dose of 150 mg/m2. Eight dogs with the porous-coated segmental bone prosthesis fixed using the EBBI principle completed the two cycles of preoperative and three cycles of postoperative multidrug chemotherapy. Clinical, radiographic, and functional gait analyses were performed every 2 weeks until the end of the experiment at 16 weeks. After sacrifice, the reconstructed femurs were subjected to radiographic, biomechanical, and histologic analyses. The results showed that the multidrug therapy had a significant effect (p  0.05) on new bone formation at the EBBI site and the ultimate torsional strength at the bone–prosthesis shoulder junction after the removal of the cement at the bone–stem interface when compared to the animal group without chemotherapy (Fig. 8). However, the newly applied grafting technique achieved better biological fixation through EBBI even with the chemotherapy when

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Figure 8 Multidrug chemotherapy effect on new bone formation measured from (A) radiographic follow-up, (B) the torsional stiffness, and (C) the maximum torque resistance attributed to EBBI alone.

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compared to our historical controls using the traditional grafting technique (particulated autogenous cortico-cancellous bone graft mixed with marrow) [100]. Therefore, the improved grafting technique was able to maintain the presence of EBBI for stem fixation even in the presence of systemic chemotherapy. However, the long-term effect of chemotherapy on EBBI maintenance needs to be studied. IV IMPROVED GRAFTING TECHNIQUE USING AUTOGENOUS BONE MARROW AUGMENTATION The concept of composite fixation, utilizing solid intramedullary stems with initial cement fixation and long-term EBBI fixation, has been developed to achieve rigid and durable fixation [21,23]. However, late complications may still occur, especially when EBBI fails to form. Therefore, closer attention has been focused on the type of tissue surrounding the implant. When successful, the bridging bone and fibrous tissue reduces bone cement interface stresses and prevents particulate debris from invading the intramedullary canal. Clinically, when EBBI is established beyond 1 year after surgery, the patient’s radiographic and functional results are good to excellent with long-term follow-up [92,109] (Fig. 9). Unfortunately, the bone graft used and its distribution remain highly variable, which could affect

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Figure 9 Follow-up examples of segmental bone replacement prostheses fixed with the EBBI concept. (A) A 32-year-old woman with chondrosarcoma in her left tibia underwent local resection and reconstruction using a single-component Ti fibermetal prosthesis fixed without bone cement in 1976. At her last follow-up in 1999, the EBBI remained practically unchanged after 23 years of active usage of the reconstructed limb. (B) A 33-year-old man with recurrent GCT which was first treated with curettage and bone grafting and fixed with a nail plate in 1974. There was recurrent GCT and the patient underwent local resection and prosthetic reconstruction in 1978. The implant stem was fractured in 1984, and the prosthesis was removed and revised using a custom Ti fibermetal segmental prosthesis and fixed using the EBBI concept. In spite of the multiple surgeries and poor bone quality, the EBBI has given successful implant long-term fixation 15 years after the revision.

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the quality of EBBI significantly. The same is true in our animal experiments, even though the amount of autogenous graft remained constant. The cause of such variation was mainly attributed to the graft preparation procedure and its fixation around the bone–prosthesis shoulder region. Accordingly, there is a strong need to develop standard and highly controllable graft preparation, distribution, and fixation methods which would significantly improve EBBI formation and maintenance. Because cancellous autogenous bone graft enhances new bone formation significantly and is relatively easily accessible, cancellous autogenous onlay bone graft first introduced by Phemister [78] is still the most widely used graft type in most clinical applications. However, cancellous bone is often resorbed quickly and provides no additional mechanical strength to the implant–bone interface. Further, achieving fixation with an even distribution of the cancellous bone graft is difficult. A new type of bone grafting with crushed cancellous bone and cortical bone strips was developed to enhance the extracortical bridging and bone ingrowth into the porous-coated prosthesis. Cancellous bone is the primary inductor of new bone formation, but the cortical bone strips are believed to enhance EBBI by contributing mechanical strength (once bridging has occurred) to the composite fixation site and acting as a scaffold for EBBI following the well-demonstrated principle of guided tissue regeneration. Six mixed-breed dogs, weighing 30 to 35 kg, were used in the experiment. All animals received a porous-coated segmental prosthesis in the middiaphyseal region of both femurs. The right limb served as an experimental side with bone graft, while the left limb received no bone graft. In the right limb, eight cortical bone strips were placed at the junction between the host bone and the prosthesis. Crushed cancellous bone was placed under and between cortical bone strips (Fig. 10). All animals were sacrificed after 12 weeks. Healing was evaluated by using gait analysis, biomechanical torsion testing, x-ray analysis, computerized CT analysis, and histomorphometry. Dynamic weight bearing showed a statistically significant decrease in ground reaction forces in the nongrafted animals between preoperative values and 4-week postoperative values. By 7 weeks the values returned to normal. On the grafted side, there were no differences between preoperative and postoperative values. The callus area was significantly larger on the grafted sides at all time points (p  0.05). The bone–prosthetic contact area was also sig-

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Figure 10 The new grafting technique for enhancing EBBI. (A) The schematic diagram of the grafting procedure. (B) The intraoperative photograph of the completed graft fixation.

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Figure 11 Comparison of stem fixation stiffness and torsional strength (contributed by EBBI alone) values between grafted side using the new technique and nongrafted control side. nificantly greater in the grafted animals. At 12 weeks, the contact area in grafted animals was 161% of the originally grafted area with uniformly distributed bone around the prosthesis shoulder. The stiffness of nongrafted specimens was only 6.2% of the stiffness of the grafted samples. The torsional strength of the interface between the bridging tissue and the prosthesis following cement removal using an innovative specimen preparation and testing technique (see Fig. 6) was four times greater for the grafted side than for the nongrafted side (Fig. 11). This experiment clearly showed that an improved cortico-cancellous bone grafting technique using onlay cortical strips augmented with cancellous autograft and marrow provided reliable and consistent new bone formation and ingrowth required for EBBI. These results were also significantly better (p  0.01) than those achieved using the traditional grafting technique in both torsional stiffness and ultimate torsional strength. V CLINICAL, RADIOGRAPHIC, AND FUNCTIONAL DATA OF EBBI PATIENTS It was found in our clinical follow-up results that if successful EBBI occurred and was maintained for the first 2 years following reconstruction, the extracortical bone remained unchanged or improved with time, thereby offering long-lasting stem fixation without aseptic loosening (Fig. 9). In addition, early formation of EBBI was closely correlated to the amount of graft used, the graft application technique, and the initial stability of the stem fixation [23,24,109]. When porous coated stems were press-fit in the medullary cavity without additional fixation, stem loosening caused by graft separation and resorption could occur. Even under successful bone ingrowth and secured stem biological fixation, severe bone resorption at the bone shoulder region due to stress bypass distally can expose the stem to high stress and fatigue fracture making stem removal extremely difficult (Fig. 12). From 1976 to 1990, 59 patients underwent local resection and reconstruction with segmental bone replacement prostheses using extracortical bone bridging and ingrowth fixation at the Mayo Clinic. Excluding those who died or were lost to follow-up less than 2 years following either the primary reconstructive surgery or revision, 43 patients were available for the present retrospective clinical, radiographic, and functional analyses. The main purpose of this study was to correlate the quality of EBBI and the long-term success of stem fixation. Because of the nature and type of the data, a multivariate logistic regression analysis was performed for the objective investigation of the clinical follow-up results. In addition, different subgroups were formed to study specific correlative relationships because of the heterogeneous nature of this database. There were 17 males and 26 females.

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Figure 12 A patient with OGS underwent local resection and reconstruction using a custom Ti fibermetal (coated only on the stem surface) proximal humeral replacement prosthesis, which was designed and prescribed by the PI in 1979. Four years later, bone resorption and stem fracture occurred which was revised with great difficulty due to bone fracture during stem removal.

The longest follow-up was 19 years, and there were 23 cases having more than 10 years of follow-up since the original surgery or the subsequent revision due to implant-related complication. There were 34 oncologic cases, and nine patients had non-neoplastic conditions. The anatomical site of replacement was in the distal femur (23 cases), proximal femur (12 cases), three each in the diaphysis of the femur and tibia, and three at the proximal humerus. There were 20 primary tumor resection and reconstruction cases, 15 revisions of an old segmental prosthesis in tumor patients (three with stem fracture and 15 with intramedullary stem aseptic loosening), and eight revisions of failed implant with massive bone loss in nontumor patients. Osteosarcoma was the most common malignant tumor (17 out of 34), followed by five chondrosarcomas, two Ewing’s sarcomas, and seven giant cell tumors, plus several miscellaneous malignant and benign tumor cases. In the nontumor group, there were two Gorham’s disease patients, one with shotgun trauma and the other was a failed total joint replacement case with massive bone loss. Titanium fibermetal–coated prostheses were used in 27 patients, and Co-Cr-Mo implants with a beaded surface were used in 16 patients. Autogenous iliac bone grafts were applied at the junction between the bone and prosthesis shoulder in 36 cases, and mixed grafts of both autogenous iliac bone and banked allogenic bone were used in six patients, with only one patient receiving allograft alone. Using the American Musculoskeletal Tumor Society (AMSTS) functional rating system [110], the overall mean functional result was rated at 69  21% at mean follow-up of 10 years and eight months. Based on the International Society of Limb Salvage (ISOLS) functional scoring system [56], the mean functional rating of the tumor group was 6820%, compared to 7322% in the nontumor group. According to the anatomical site, the mean rating of functional results was best in the tibial diaphysis (9027%) at a mean follow-up of 171 months, followed by femoral diaphysis (8927%) at 149 months, proximal humerus

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(7021%) at 118 months, proximal femur and hip joint (5918%) at 120 months, distal femur and knee joint (5817%) at 118 months. The complications that needed reoperation were aseptic loosening in six patients (five incidences occurred without initial cement fixation, and two loosenings occurred in the same patient)—five stem fractures in three cases (Fig. 13) and one taper joint fracture (Fig. 14)—three insufficient extracortical bone bridging (all of which achieved EBBI after regrafting), two infections, and two dislocations (in the same patient). There were three late loose acetabular components in proximal femoral and hip replacement patients and one late tibial component loosening in a distal femur and knee replacement. The measurement of EBBI was performed using anteroposterior and mediolateral radiographs containing the bone–prosthesis body junction area where bone grafting was applied. Bone formation over the thickness of 1.5 mm covering 2.5 cm in length over the porous-coated shoulder region of the implant was considered 100% EBBI coverage. In each stem, there are four critical zones to be measured: the medial, the lateral, the anterior, and the posterior zones. The mean value of EBBI, which is the average of the measurements at the four critical zones, was used in the present analysis. Each case was evaluated for its 2-year mean EBBI value and the last follow-up EBBI value for comparison. Due to the incomplete radiographic data in each case, the initial time periods varied from 1 to 4 years for the present study. Using the paired Student’s t test, we found that after initial successful EBBI formation (we used 25% of mean EBBI coverage as a subjective reference), the difference in the amount of EBBI between the initial and final measurements was not significantly different (p  0.0001). There were several cases which showed apparent improvement of EBBI with time. The data generated from the present study were affected by the quality of the radiographs and whether they had adequately covered the critical zones to be analyzed.

Figure 13 A 16-year-old girl with OGS in her right distal femur and knee underwent local extracapsular resection and knee fusion using a fully Ti fibermetal coated modular implant fixed without bone cement. The proximal stem fractured 10 months postoperatively, but healed through EBBI without any symptoms. The distal stem was also fractured during a car accident 8 years after the initial surgery. The patient was treated conservatively and the fracture healed through EBBI also uneventfully.

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Figure 14 The only case with fractured Morse taper joint in this 14-year-old boy with OGS in his right proximal humerus, which was resected and replaced by a custom Ti fibermetal modular implant with cement fixation of the extremely short distal stem (2 in. long). The taper joint fracture did not produce any clinical symptoms. There were no patients with late stem loosening once successful EBBI was established. Using the multivariate logistic regression model, the important risk factors for prosthesis loosening were the type of implant used (odds ratio: 16.4, p  0.04) and the amount of EBBI formation (odds ratio of 10% increment in percentage of coverage: 0.77, p  0.10). The Ti fibermetal prosthesis had a higher incidence of stem loosening, since the majority of the prostheses of this type had porous-coated stems and were press-fit into the intramedullary cavity. Additional important information can be derived from the current long-term follow-up clinical data. In all of the six cases involving the use of allograft, either in combination with autogenous bone graft or applied alone, the EBBI quality was superior, with no stem loosening. There were no late complications in the stems with successful EBBI formation. In studying serial radiographs, initial graft amount and its distribution seemed to influence the quality of EBBI. In addition, the current retrospective study was further limited by the exclusion of all the early aseptic loosening cases (less than 2 years) which involved noncemented fixation and no or poor grafting technique. We do feel strongly that extracortical bone bridging and ingrowth will provide a successful clinical outcome in the reconstruction of large bony defects resulting from primary tumor resection, failed tumor prosthesis, and failed joint arthroplasty with extensive bone loss. However, additional studies and further refinements will be necessary to assure the effectiveness and reliability of this concept as proposed in the present revised application. VI METAL ION RELEASE IN PATIENT URINE AND BLOOD SERUM In segmental bone and joint prosthetic replacements, a porous coating is used either on the intramedullary stem or on the body of the implant in order to achieve biological fixation.

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This type of prosthesis is usually much larger than a regular joint replacement implant and, combined with added porous coating, provides expanded metallic surface exposure to tissue and body fluids. The objective of this study was to quantify the trace metal ions in patients with two types of implant metals with a follow-up of at least 2 years. The normal values for Ti, Al, V, Co, Cr, and Mo were established in 12 women and 12 men with matching age and body weight with the patient population. The patients were divided into two groups, a Ti fibermetal segmental prosthesis group (n  17) and a Co-Cr-Mo porouscoated megaprosthesis group (n  10). Patients with Ti implants were followed longer (mean of 6.2 years) than those with Co-Cr-Mo prostheses (mean of 4.5 years). The porous area ranged from 46 to 514 cm2. Seven out of 27 patients were operated on for failed arthroplasty with massive bone loss, and the remaining 20 patients had reconstruction following tumor resection. Urine samples and blood (2 to 3 mL) serum specimens were collected from the normal subjects and the patients following a strict protocol using special containers to prevent contamination. Standard metal tracing detection techniques (furnace atomic absorption) were used by the Toxicology Lab at Mayo under careful environmental controls. In analyzing Ti concentration, electrodeless discharge lamps (EDLs) under controlled ionization were used. Statistical correlations using ANOVA were performed to relate porous surface size, follow-up length, presence or absence of taper lock in modular design, type of implant materials, and pathological conditions with the metal ion concentration. The values of metal ion release in urine and blood serum were all insignificant in the normal subjects. In patients with prosthetic implants made of Ti alloy (Ti-6A1-4V), the Ti level was significantly higher (p  0.001) than that of the normal reference found either in urine or in blood serum. Elevation of Al was small, except in urine, and V concentration was insignificant. In the Co-Cr-Mo group, the elevation of Mo was significant (p  0.005). Co and Cr levels were comparable to the normal standard. In special cases involving stem loosening or prosthetic fracture, the related metal ion release can increase 50- to 200-fold. In a patient with a metal hinge prosthesis, the metal ion concentration in the nearby fluid and tissue was enormous, with concomitant bone resorption. There appears to be an association between ion release amount and porous area size, but such relationships were weak among implants made of Co-Cr-Mo material. Using the normal level as the base, Ti alloy implants appear to have high ion concentration. There was no relationship between the amount of ion release with either follow-up time or pathology. Mechanical failure was found to be the main factor for local and systemic metal ion concentration. Under normal situations, metal ion release does not seem to present major concerns in megaprostheses used in limb salvage surgery. However, when a taper lock is present, loosening or fracture can cause significant metal ion release through wear and corrosion. If a fibrous tissue capsule is formed around the taper lock through an adjacent porous surface, it can serve as a metal particle/ion barrier to prevent further migration of these particles. Metal-to-metal hinge interfaces should not be used since the metal concentration in local tissues will inevitably lead to bone resorption. Porous-coated large segment prostheses have evolved as the dominant reconstruction device following major bone loss and articular surface loss. Further improvements in durability can be achieved through implant design and fixation refinements. VII DIRECT TENDON REATTACHMENT TO METALLIC PROSTHETIC ANCHOR A series of in vivo experiments utilizing a canine model were performed to develop the methodology for attaching a tendon to an enhanced tendon anchor (ETA) prosthesis (Fig. 15) [13,14,42,43,53]. The prosthesis surface was composed of a titanium mesh sintered to

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Figure 15 The enhanced tendon anchor (ETA) device. a supporting substructure, with a void space that permitted ingrowth of bone and soft tissue. Initial tendon fixation strength was provided by a small titanium plate stabilized by two set screws with interlocking spikes. For each experiment, the fixation strength, clinical function of the tendon, and the morphological changes at the attachment site were examined. The ETA prosthesis/bone graft composite was tested under in vitro condition for its fixation strength using a canine shoulder complex. The tendon compressive force was calibrated using a torque wrench. The initial compressive force was found to relax with time up to 15 min. The tendon pull-out strength was directly related to the tightening torque of the fixation screws [40]. To avoid damage to the tendon and the bone graft, a tightening torque of 0.15–0.2 N-m was found to be optimal for the initial fixation of the tendon to the metallic implant (the ETA prosthesis) with the bone graft sandwiched between. The tendon pull-out strength was found to be 21–32% of the intact insert strength. Three in vivo experiments were studied based on different tendon attachment methods (Fig. 16). In the first experiment involving 12 canines, both supraspinatus tendons were attached to a titanium prosthesis. In one limb, the bone block underlying the tendon inser-

Figure 16 Schematic diagrams showing the four tendon attachment methods. BB, fixation with intact bone block; BPG, fixation with bone plate sandwiched between ETA and tendon with bone marrow augmentation; BP, fixation alone without marrow augmentation; TA, direct tendon attachment to ETA device.

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tion was preserved and attached to the metal prosthesis. In the contralateral limb the tendon was attached directly to the prosthesis. The tensile fixation strength was evaluated 16 weeks after surgery. Six cadaveric canine shoulders were similarly tested to serve as intact normal controls. The tendon–bone block attachment was significantly stronger than the direct tendon attachment (p  0.05), but not significantly different from the normal control (Fig. 17). Weight bearing was greater on the limbs with a direct tendon attachment at both 3 and 6 weeks postoperatively. Both front legs showed a trend for increased weight bearing with time. Bone ingrowth into the prosthesis was consistently found for the tendon–bone block attachment, while fibrous tissue ingrowth was found in the direct tendon attachments (Fig. 18). When a bone block was preserved, the strength and stiffness were comparable with those of a normal tendon insertion (Fig. 17). In a second study, the supraspinatus tendon was attached to the ETA device with an interpositional autogenous cancellous bone plate between the tendon and the mesh surface of the prosthesis. In 11 dogs, the right supraspinatus tendon was attached to the ETA prosthesis. An autogenous cancellous bone plate (12 mm  20 mm  2 mm) was pressed upon the fixation spikes against the mesh surface (Fig. 16). The tendon was then placed over the graft and held in place by the fixation plate. The contralateral limb was left unoperated. All dogs used the operated limb immediately after surgery. An initial decrease in weight bearing was observed 3 weeks after surgery, followed by a significant increase in weight bearing by 6 weeks (p0.0001). The interpositional bone graft remained visible on radiographs until 12 weeks after implantation. Early changes included loss of distinct graft margins, and late changes were characterized by partial to total loss of graft at the surface of the implant. The stiffness and strength of the reconstructed side were significantly less than the strength and stiffness of the intact contralateral that was most active between 8 and 16 weeks after implantation. Very little new bone formation was observed at any time. Soft tissue was highly organized with prominent anchoring fibers. Histologic sections revealed progressive osteoclastic resorption of the graft ingrown within the mesh after 16 weeks [54]. In a third study, the effects of using an interpositional autogenous cancellous bone plate and bone marrow grafting on soft tissue reattachment to a metallic prosthetic implant was studied using the canine shoulder model in nine dogs [55]. Supraspinatus tendon attachments were reattached to the mesh surface of the ETA prosthesis. An interpositional autogenous cancellous bone graft harvested from the removed tubercle of the humerus was fastened to the prosthesis between the mesh and the tendon (Fig. 19). Additional autogenous

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Figure 17 Mechanical strength comparison among the four tendon attachment methods—(A) failure strength; (B) fixation stiffness. BB, fixation with intact bone block; BPG, fixation with bone plate sandwiched between ETA and tendon with bone marrow augmentation; BP, fixation alone without marrow augmentation; TA, direct tendon attachment to ETA device.

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Figure 18 (A and B) Contact microradiograms showing bone ingrowth into the porous space of the ETA device when tendon attachment was achieved through intact bone block. (C) Residual bone left on and within the mesh after mechanical tensile test to failure.

bone graft and bone marrow was supplemented beneath the mesh of the implant and around the interpositional bone graft and reattached tendon. Gait analysis and radiological analysis were performed every 3 weeks after the surgery. Histological analysis and mechanical testing in tension was conducted 16 weeks after the surgery. Weight bearing increased significantly by 6 weeks. At 16 weeks after surgery weight bearing recovered to 90.3% of the preoperative level, which was significantly higher than weight bearing for fixation using the bone plate alone. Muscle volume was 85.5% of contralateral side, which was also higher

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Figure 19 (A) Interpositional autogenous cancellous bone plate with bone marrow augmentation and (B) the contact microradiogram showing cancellous bone plate graft surviving the application period with limited bone ingrowth into the porous mesh at 16 weeks after surgery.

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Figure 20 (A) The regeneration of the four tissue morphology zones for tendon attachment using the ETA device combined with (B) autogenous bone plate augmented with cancellous bone and marrow. T, tendon tissue; U, uncalcified fibrocartilage; C, calcified fibrocartilage; B, the subchondral cancellous bone. than fixation with the bone plate alone. Ectopic calcification was observed around the prosthesis and its area increased by 63% at 15 weeks (p  0.06). The mean tensile stiffness of the reconstruction was 73.0 N/mm, and the mean ultimate tensile strength was 726.0 N. These values were significantly less than the stiffness and strength of the intact contralateral side (55.2 and 38.6%, respectively) and the reconstruction with an intact bone block, but significantly larger than those of direct tendon attachment without the interpositional bone graft or with the interpositional bone plate and without sufficient bone marrow. The bone plate with marrow augmentation achieved significant improvement through the formation of a pseudosubchondral bone plate similar to that fixed with the intact bone block. The autogenous bone marrow initiated an osteoinductive reaction evidenced by the radiographic and histological results. New bone formation was found in the bone plate, at the mesh surface, and in the tendon. This newly formed tendon/bone unit acted as a macroscopic anchor to allow direct load transmission between the supraspinatus tendon and the ETA under physiological loading. Morphologically, this process appeared to reproduce the tendon fiber zone, the fibrocartilage zone, the calcified fibrocartilage zone, and the subchondral cancellous bone (Fig. 20). Regeneration of the tendon/bone unit using an autogenous bone plate and marrow grafting over the ETA prosthesis provided a novel and effective means of the reattachment of the soft tissue. In osteochondral allograft and composite reconstructions involving the proximal femur, humerus, and tibia, soft tissue reconstruction could be achieved through macroscopic and microscopic anchorage as demonstrated in this study. VIII BIOLOGIC ENHANCEMENT FOR GUIDED TISSUE REGENERATION To improve graft fixation and EBBI formation, homologous fibrin adhesive was applied to prevent graft migration and to improve osteogenic induction potential, but the enhancement on EBBI was insignificant [84]. In an attempt to find an adequate bone graft substitute to form EBBI, demineralized bone matrix and hydroxyapatite coating on the porous surface were studied, but the results were uniformly inferior to that achieved using the autograft [33,104]. Bone morphogenetic proteins have become the most effective and reliable bone induction agent for heterotopic bone formation, bone defect repair, and fracture healing augmentation [81,97,98]. From our recent long-term patient follow-up data and the results of our animal experiment with autogenous cortical and cancellous graft, we have formu-

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lated our current working hypotheses to investigate an autogenous bone graft substitute and osteogenic enhancing factors to establish EBBI using rhBMP with collagen-I sponge as a carrier combined with allograft cortical strips as a structural scaffold. Bone morphogenetic proteins were isolated from bone initially by acidification and subsequent chromatography [45,85,99]. These proteins are the most potent growth factors in both in vitro and in vivo assays of bone formation [80,106]. In addition, one of these growth factors has been used in a clinical tibial nonunion trial with approximately equivalent results to the positive control using autogenous bone graft [58]. The effect of these growth factors has not been studied as an autogenous bone graft substitute, but one specific growth factor (rhBMP-7) is currently being investigated in our laboratory as an adjunct to improve segmental allograft osteosynthesis in a mid-femoral defect model. The rhBMPaugmented allograft bone may be the ideal composite to induce bone formation and thus bridge the porous prosthesis surface and host bone through extracortical bone formation. If this new grafting technique of combining allograft with rhBMP is equivalent to or significantly better than the autogenous graft, the EBBI principle can be more effectively and reliably utilized to improve segmental bone replacement prosthesis fixation. Such technique has several advantages over the traditional autograft technique since (1) autogenous graft material is always limited in patients undergoing limb salvage surgery; (2) graft harvesting may produce significant donor site morbidity and occasional complications; and (3) the current grafting technique lacks uniformity due to graft size, volume variation, and distribution. Specifically, the rhBMP will be delivered using a collagen-I sponge at a concentration proven to be effective to stimulate osteogenesis and bone formation at the junction between the porous-coated prosthesis and the host bone cortex strutted with allograft cortical strips. Recently, clinical studies using these two recombinant human bone morphogenetic proteins, rhBMP-2 and rhBMP-7, for bone regeneration have been reported [5,40,44]. Since the formation of extracortical bone bridging with the autogenous bone graft included appositional, intramembranous, and endochondral ossification mechanisms, the role of both rhBMP-2 and rhBMP-7 need to be investigated and compared in the same experimental settings. These growth factors have demonstrated significant variation in their bone regeneration characteristics [4,48,72,75,86]. These existing data provided the rational basis to select the rhBMP-2 and rhBMP-7 for clinical trial once proven to be effective and highly reliable in well-controlled animal experiments. IX FUTURE DEVELOPMENT IN SEGMENTAL BONE AND JOINT REPLACEMENT To further improve the grafting technique for EBBI induction, we have developed a unique device, the Makisu device, to incorporate the allograft strips, the rhBMP-impregnated collagen-I sponge strips, and the nonresorbable suture in a totally flexible and easy-to-adjust form to accommodate bone size and bone–prosthesis diameter mismatch, while achieving secure fixation at the desirable location (Fig. 21). The word Makisu implies a flexible mat made of bamboo strips and connected by strings with proper space between strips to be used in forming Japanese “roll sushi.” The current device will contain the collagen-I sponge strip to make up the space between the allograft strips. Since the sponge is soft and flexible, the unique feature of Makisu will be retained to allow the application of the device around the prosthesis–bone junction where EBBI is to be formed. The placement of the rhBMP in the collagen-I sponge will be achieved by injecting the protein in the predeter-

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Figure 21 The new grafting technique for EBBI using guided tissue regeneration principles and tissue engineering concept. The proposed new grafting technique using the allograft and rhBMP-impregnated collagen-I sponge composite construct, called the Makisu device, to be applied around the implant–bone junction to enhance EBBI induction. The rhBMP at known concentration and volume will be applied through a syringe onto the sponge just before the Makisu device is surgically placed around the desirable location and tied down by the sutures.

mined concentration and volume with a syringe. After the surgical field and the reconstruction sites are thoroughly irrigated with saline solution, the Makisu device will be applied around the locations where EBBI is strategically needed for composite fixation of the segmental bone replacement prosthesis. We strongly believe that the present research will be necessary to make the EBBI principle easier to accomplish with improved quality, assuring better segmental bone replacement prosthesis fixation. Once this principle and technique have been further refined, they can be adopted by all segmental bone replacement prosthesis systems to improve their clinical outcome and durability. The present application is specifically aimed to seek optimal effectiveness in the proposed biological fixation concept to bone in segmental bone/joint replacement prosthesis. This reconstructive technique and related implant design principles will also benefit a large patient population with massive bone defects secondary to trauma, metabolic bone diseases, and failed total joint replacement arthroplasty [67,90]. If successful, this new technique will have a profound impact on healing bone defects that are juxtaposed to metal fixation devices, principally, intramedullary rods and prosthetic stems. To achieve tendon attachment to metallic implant, the use of appropriate connective tissue growth factors should be considered since the normal tendon attachment morphology has four distinctive tissue zones with drastically different structures and properties. The subchondral cancellous bone and the uncalcified fibrocartilage may be relatively easy to

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generate using either the guided tissue regeneration principles or the tissue engineering concept. However, the calcified fibrocartilage would require functional loading and activity to transform and remodel the existing tissue to complete the regeneration process. Therefore, long-term follow-up is required to validate the currently proposed tissue regeneration concept for tendon reattachment to metallic implant for its intended physiological and biomechanical functions. The entire concept of segmental bone and joint replacement using biologically augmented prosthesis was developed following the general principles of guided tissue regeneration. Under ideal conditions, such a method may be sufficient to produce the desirable tissue for the functional requirements relying only on a patient’s own biological potential. Unfortunately, autogenous graft materials in the body are always limited in supply and the undesirable local and systemic environment associated with the patient’s prior medical condition and his/her concurrent adjuvant treatment, which would further jeopardize cell recruitment and transformation for tissue formation. These deficiencies necessitate the application of the tissue engineering concept where the biological factors will be delivered or stimulated from external sources. Such enhancement factors can be applied in different forms of drugs, specific lines of cells, or even through biophysical stimulation. The biophysical stimulation could be delivered noninvasively using ultrasound, DC or AC signals, electromagnetic fields of different frequency, power, and waveform, which provide the exciting possibility of achieving noninvasive tissue engineering!

Figure 22 Segmental bone and joint replacement prosthesis for the proximal humerus and shoulder using the EBBI concept for implant fixation to bone and the ETA device for direct tendon reattachment.

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Finally, it is important to realize that tissue engineering has its limitations, especially related to musculoskeletal systems. In addition to the functional requirements which could only be established through a remodeling process, tissue engineering under its current development has not properly addressed the needs at the organ and system levels. Segmental bone and joint replacement involves different types of tissue and they must satisfy repetitive loading demands in our activities of daily living and recreation. Much of the material, structural, anatomical, physiological, and biomechanical knowledge related to out musculoskeletal system function is still unknown, it would be farfetched for us to imagine that the current concept of tissue engineering will one day lead us to regenerate a whole functional limb. With the methods introduced in this chapter combined with guided tissue regeneration principles and the tissue engineering concept, it is conceivable that limb salvage surgery can be significantly improved by adopting EBBI in prosthesis design and implant fixation and the ETA device for soft tissue attachment to benefit many patients (Fig. 22). ACKNOWLEDGMENTS This work was supported by NIH Grant CA 23751, awarded by NCI, NIH, DHHS from 1978 to 1999. Many fellows and coworkers of the authors both at the Mayo Clinic and at Johns Hopkins University contributed greatly to the investigations summarized. The authors also wish to express their thanks to the orthopedic industry which transformed the outcome of this research into prosthetic implants to benefit patients selected to undergo limb salvage procedures after tumor resection or failed joint replacement surgery with massive bone loss. Finally, we wish to pay tribute to those in the field who supported and encouraged the ideas presented in this chapter. REFERENCES 1. Aro H. T., Aho A. J. 1993. Clinical use of allografts. Ann. Med. 2:403–412. 2. Alman B. A., De Bari A., Krajbich J. 1995. Massive allografts in the treatment of osteosarcoma and ewing sarcoma in children and adolescents. J. Bone Joint Surg. 77A:54–64. 3. Bacci G., Picci P., Ferrari S., Ruggieri P., Casadei P., Tienghi A., Brach del Prever A., Gherlinzoni F., Mercuri M., Monti C. 1993. Primary chemotherapy and delayed surgery for nonmetastatic osteosarcoma of the extremities. Results in 164 patients preoperatively treated with high doses of methotrexate followed by ciplatin and doxorubicin. Cancer 72:3227–3238. 4. Boden S. D., McCuaig K., Hair G., Racine M., Titus L., Wozney J. M., Nanes M. S. 1996. Differential effects and glucocorticoid potentiation of bone morphogenetic protein action during rat osteoblast differentiation in vitro. Endocrinology 137:3401–3407. 5. Boden S. D., Zdeblick T. A., Sandhu H. S., Heim S. E. 2000. The use of rhBMP-2 in interbody fusion cages. Definitive evidence of osteoinduction in humans: a preliminary report. Spine 25(3):376–381. 6. Bell R. S., Davis A., Allan D. G., Langer F., Czitrom A. A., Gross A. E. 1994. Fresh osteochondral allografts for advanced giant cell tumors at the knee. J. Arthrop. 9:603–609. 7. Bradish C. F., Kemp H. B. S., Scales J. T., Wilson J. N. 1987. Distal femoral replacement by custom-made prostheses. J. Bone Joint Surg. 69B:276–284. 8. Brien E. W., Terek R. M., Healey J. H., Lane J. M. 1994. Allograft reconstruction after proximal tibial resection for bone tumors. An analysis of function and outcome comparing allograft and prosthetic reconstruction. Clin. Orthop. 303:116–127. 9. Campanna R., Van Horn J. R., Biagini R., Ruggieri P., Bettelli G., Sola G., Campanacci M.

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91. Sim F. H., Beauchamp C. P., Chao E. Y. S. 1987. Reconstruction of musculoskeletal defects about the knee for tumor. Clin. Orthop. 221:188–201. 92. Sim F. H., Frassica F. J., Chao E. Y. S. 1995. Orthopaedic management using new devices and prostheses. Clin. Orthop. 312:160–172. 93. Simon M. A., Aschiliman M. A., Thomas N., Mankin H. J. 1986. Limb-salvage treatment versus amputation of the distal end of the femur. J. Bone Joint Surg. 68A:1331–1337. 94. Simon M. A. 1986. Current concepts review: limb salvage for osteosarcoma. J. Bone Joint Surg. 68:1331. 95. Springfield D. 1996. Autograft reconstruction. Orthop. Clin. N. Am. 27:483–492. 96. Stevenson S., Emery S. E., Goldberg V. M. 1996. Factors affecting bone graft incorporation. Clin. Orthop. 324:66–74. 97. Stevenson S. 1998. Enhancement of fracture healing with autogenous and allogeneic bone grafts. Clin. Orthop. 355S:S239–S246. 98. Trippel S. B. 1998. Potential role of insulinlike growth factors in fracture healing. Clin. Orthop. 355S:S301–S313. 99. Urist M. R. 1965. Bone formation by autoinduction. Science 150:893–899. 100. Virolainen P. H., Inoue N., Nagao M., Ohnishi I., Frassica F. J., Chao E. Y. S. 1999. Effect of bone graft on extracortical bone and capsule formation over porous segmental prosthesis. J. Bone Joint Surg. 81A:493–499. 101. Wanebo H. J., Temple W. J., Popp M. B., Constable W., Aron B., Cunningham S. L. 1995. Preoperative regional therapy for extremity sarcoma. A tricenter update. Cancer 75:2299– 2306. 102. Ward W. G., Johnston K. S., Dorey F. J., Eckardt J. J. 1993. Extramedullary porous coating to prevent diaphyseal osteolysis and radiolucent lines around proximal tibial replacements. A preliminary report. J. Bone Joint Surg. 75A:976–986. 103. Warren S. B., Pelker R. P., Friedlaender G. E. 1985. Effects of short-term cyclosporin-a on biomechanical properties of intact and fractured bone in the rat. J. Orthop. Res. 3:96–100. 104. Wippermann B. W., Hsu R. W. W., Chao E. Y. S., Sim F. H. 1991. Comparison of autogenous cortical graft and demineralized allogenic bone matrix (DABM) in the fixation of segmental prosthesis. In: Limb Salvage—Major Reconstructions in Oncologic and Nontumoral Conditions, Langlais F., Tomeno B., Eds. Springer-Verlag: Berlin, pp. 335–343. 105. Wood M. B., Cooney III W. P., Irons Jr. G. B. 1985. Free vascularized fibula graft transplantation: skeletal indications and results. Mayo Clin. Proc. 60:729–734. 106. Wozney J. M., Rosen V., Celeste A. J., Mitsock L. M., Whitters R. W., Kriz R. W., Hewick R. M., Wang W. A. 1988. Novel regulators of bone formation: molecular clones and activities. Science 242:1528–1534. 107. Young D. R., Virolainen P., Inoue N., Frassica F. J., Chao E. Y. S. 1997. The short term effect of cisplatin chemotherapy on bone turnover. J. Bone Miner. Res. 12:1874–1882. 108. Young D. R., Shih L.-Y., Rock M. G., Frassica F. J., Virolainen P., Chao E. Y. S. 1997. The effect of cisplatin chemotherapy on extracortical capsule formation in canine segmental replacement. J. Orthop. Res. 15:773–779. 109. Shin D. S., Choong P. F., Chao E. Y., Sim F. H. 2000. Large tumor endoprostheses and extracortical bone-bridging: twenty-eight patients followed 10–20 years. Acta Orthop Scand. 71(3):305–311. 110. Enneking W. F., Dunham W., Gebhardt M. C., Malawer M., Pritchard D. J.: A system for functional evaluation of reconstructive procedure after surgical treatment of tumors of the musculoskeletal system. Clin. Orth. Rel. Res. 286:241–246, 1993.

20 Injectable Calcium Phosphate Cements for Repair of Bone Defects Shigeo Niwa and Racquel Z. LeGeros Aichi Medical University, Aichi, Japan and New York University College of Dentistry, New York, New York

I INTRODUCTION The earliest application of a calcium phosphate reagent labeled as “triple calcium phosphate” (of unknown composition) in a successful repair of bone defect was reported by Albee in 1920 [1]. More than 50 years later, Nery et al. reported the first application in a surgically created periodontal defect of a synthetic calcium phosphate preparation they described as a tricalcium phosphate (TCP) [2]. The “TCP” used by Nery was later analyzed, using x-ray diffraction, as consisting of a mixture of calcium hydroxyapatite (HA) and beta-tricalcium phosphate (-TCP) with an HA/-TCP ratio of 80/20 [3]. Since then, such materials have been more appropriately described as biphasic calcium phosphates (BCP) with varying HA/-TCP ratios [4–8]. The commercialization of synthetic HA as a bone graft substitute material for dental and orthopedic applications was pioneered by Jarcho [9], Aoki [10], and deGroot [11]. To date, synthetic calcium phosphates used as bone graft materials include HA, -TCP, BCP, coralline HA (prepared by hydrothermal conversion of coral, a calcium carbonate), bovine bone–derived HA (treated and sintered bovine bone), and bovine bone–derived apatite (unsintered). The rationale for developing synthetic calcium phosphates as synthetic bone graft materials is their similarity in composition to that of the bone mineral. The bone mineral was idealized as HA, Ca10(PO4)6(OH)2 [12,13]. However, based on a combination of analytical methods (x-ray diffraction, infrared, and chemical analyses) and compared with synthetic carbonate-containing apatites prepared at 37 to 100°C), the bone apatite, like other biological apatites (the mineral phases of enamel, dentin and bone, and other pathological calcifications), was determined to be a carbonate hydroxyapatite ap385

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Figure 1 X-ray diffraction patterns of (A) synthetic hydroxyapatite, HA, (B) bone mineral, and (C) synthetic carbonate containing apatite, CHA.

proximated by the formula (Ca,Na,Mg)10(PO4,HPO4,CO3)6(OH)2, with the carbonate (CO3) as a minor but important substituent (principally for the PO4 group) [14,15]. Biological apatites form under physiological temperature (37°C). On the other hand, pure hydroxyapatite, Ca10(PO4)6(OH)2, is obtained by precipitation under very basic conditions and subsequently sintered at approximately 1100°C [9–11,16]. Thus, besides the composition (Fig. 1), bone apatite and synthetic HA differ in several properties such as crystallinity, reflecting crystal size (Fig. 2) and solubility (Fig. 3). Because the carbonate hydroxyapatite crystals of

Figure 2 Fourier transform infrared spectra of (A) synthetic HA, (B) bone mineral showing the presence of carbonate (CO) in bone apatite, and (C) synthetic carbonate containing apatite, CHA.

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Figure 3 Comparative dissolution of (A) HA and (B) bone mineral in acidic buffer (0.1M Kac, pH 6, 37°C, powder-to-liquid ratio 25 mg:100 mL).

the bone mineral forms under physiological temperature (37°C), they are much smaller than the HA crystals [14,16]. The combination of nanocrystal size and incorporation of carbonate in the apatite make the bone mineral much more soluble than synthetic HA [14,17]. Synthetic calcium phosphates are characterized as biocompatible, bioactive, and osteoconductive [3,16–21]. In the form of particulate or blocks and as coatings on orthopedic and dental implants, calcium phosphates have been used in several dental and medical applications, e.g., repair of bone defects, alveolar augmentation, maxillofacial reconstruction, ear implants, spine fusion, etc. [3–11,19–23]. However, in spite of their desirable properties, this generation of calcium phosphate biomaterials had some limitations: slow bioresorption, which does not allow it to be replaced by the newly forming bone, inability to be used in irregular-shaped defects, and difficulty in handling and delivery for some special clinical applications. It was apparent that there was a need for materials which more closely approximate the properties (e.g., crystallinity and bioresorbability) of the bone mineral; the development of calcium phosphate cement was an attempt to meet this need. II CALCIUM PHOSPHATE CEMENT FORMULATIONS AND PROPERTIES The concept and potential advantages of apatitic or calcium phosphate cement (CPC) as possible restorative material was first introduced by LeGeros et al. in 1982 [24,25]. The early preliminary cement formulation was based on mixing calcium-deficient apatite (HAp) and calcium hydroxide with dilute phosphoric acid solution [24]. In 1987, Brown and Chow reported the first self-hardening CPC resulting from the reaction of tetracalcium phosphate (TTCP) and dicalcium phosphate anhydrous (DCPA) with distilled water or sodium phosphate solution [25,26]. Other CPC formulations have been developed in recent years [28–36] and some have become commercially available as shown in Table 1 [36]. Calcium phosphate cement systems consist of a solid and a liquid component. The solid component may be made up of one calcium phosphate compound or a mixture of calcium phosphate compounds or a mixture of calcium phosphate and other nonphosphatic

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Table 1 Calcium Phosphate Cements (Commercial) Cement name -BSM Biobone Embarc Norian SRS, Norian CRS Fracture grout Bonesource Cementek Biocement D Biopex a

Company ETEX Merck GmgH Lorenz Surgical Norian (Synthes)

Leibinger, Orthofix Teknimed Merck GmbH Mitsubishi Materials

Solid component

End producta

ACP, DCPD

HAp

-TCP, CaCO3, MCPM

CHA

TTCP, CaCO3, H3PO4 TTCP, DCPA -TCP, TTCP, MCPM -TCP, DCPA, CaCO3, HAp -TCP, DCPA, CaCO3, HAp

HAp HAp HAp CHA CHA

End product is the product obtained after the cement sets in vitro. Except for Biopex end product, which was confirmed by studies reported in this paper, the end products of other cements are quoted from reports in the literature.

calcium compounds. The calcium phosphate compounds include amorphous calcium phosphate (ACP), Cax(PO4)yH2O; monocalcium phosphate monohydrate (MCPM), CaH4(PO4)2H2O; dicalcium phosphate dihydrate (DCPD), CaHPO42H2O; dicalcium phosphate anhydrous, CaHPO4; octacalcium phosphate (OCP), Ca8H2(PO4)65H2O; calcium-deficient apatite, Ca10x(PO4,HPO4)6(OH)2; alpha- or beta-tricalcium phosphate (or -TCP), Ca3(PO4)2; and tetracalcium phosphate, Ca4P2O9. These calcium phosphate compounds differ in their Ca/P molar ratio, solubility, and propensity to transform or hydrolyze to apatite or carbonate apatite in vitro or in vivo, depending on the composition and pH of the environment [16,25,37]. The nonphosphatic calcium compounds include calcium oxide, CaO; calcium hydroxide, Ca(OH) 2; calcium carbonate (CC), CaCO3; and calcium sulfate hemihydrate (CSH), CaSO40.5H2O). The liquid component may be distilled water, saline solution, sodium phosphate solution, dilute phosphoric acid, dilute organic acids (e.g., succinic), sodium alginate, or sodium chrondroitin sulfate, etc. The setting reaction product(s) or product obtained after setting is determined by the composition of the solid and the composition and the pH of the liquid phase. The setting time, which can range from 5 to 60 min, is determined by the composition of the solid and liquid components, the powder-to-liquid ratio, the proportion (e.g., TTCP/DCPA ratio) of the powder components, and the particle size of the solid component(s). Apatitic calcium phosphate or carbonatecontaining apatite (CHA) (with crystallinity similar to that of bone) can form before implantation when the cement sets or can result from the in vivo hydrolysis of the nonapatitic calcium phosphate (e.g., DCPD) after implantation. Compared to HA, -TCP CPC, which is usually only available in the form of particulates or blocks, has the following desirable useful properties and decided advantages: malleability, which allows it to adapt to the size and shape of the defect, and high bioresorbability, which allows it to be replaced by bone. A Injectable Calcium Phosphate Cement The ease of handling and delivery of CPC was greatly improved by the development of injectable calcium phosphate cements which opened up new avenues of treatment.

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The commercial and experimental injectable calcium phosphate cements share similar properties and potential dental and medical applications. This chapter describes highlights of the results of intensive studies on the properties and clinical applications of one injectable calcium phosphate cement, Biopex® (Mitsubishi Co.), which was developed in 1989. 1 Composition The powder component of Biopex consists of 75 %wt -TCP, 18 %wt TTCP, 5 %wt DCPD, and 2 %wt HA. The liquid component consists of 5% sodium chondroitin sulfate, 12% sodium succinate, and 83% distilled water. Ratio of powder to liquid (P/L) is 1.0:3.2. The setting time, depending on the P/L ratio, is 6 to 10 min at 37°C [38]. 2 In Vitro and in Vivo Transformation When soaked in saline solution, the CPC transforms to carbonate apatite as early as after 7 days, as shown by FTIR analysis. Similar transformation was observed after implantation in vivo [39], as shown in Figs. 4 and 5. 3 Histocompatibility of CPC Microscopic observation of the CPC injected into the medullary cavity of the rabbit femur showed newly formed and well-developed bone tissue that also invaded into the CPC, 4 weeks after injection. No fibrous tissue was observed between the bone and the CBC [39], shown in Fig. 6a,b. Scanning electron microscopic (SEM) observation revealed macrophages surrounding the CPC at 3 days after injection into the medullary cavity of the rabbit femur. At 2 weeks after injection, the periphery of the CPC was covered by many osteoblasts. In addition, newly formed bone was well developed [39], as shown in Fig. 7a,b. Sugimoto et al. [40] described that small blood vessels extended into the CPC injected into the medullary cavity of a rabbit femur at 8 weeks (Fig. 8). This was an important observation since vascularization was not observed with implanted HA. 4 Structure/Function Relationship The mechanical strength of materials for filling bone defect is an important factor in clinical applications. Injectable bone cements are not only easy to fit in bone defect with irregular shape (e.g., in the case of osteoporotic cancellous bone with enlarged porous cavities), but also increase the mechanical strength of the cancellous bone. The compression strength of the CPC in this report was about 80 MPa in vitro and in vivo by 1 week after mixing, and then kept a steady state for 4 weeks [40], as shown in Fig. 9. The efficiency of the CPC in fracture treatment was investigated by Zhang et al. [41] using 50 Chinese mountain sheep in which the femoral neck was osteotomized to serve as an experimental model of femoral neck fracture. The osteotomized femoral necks were fixed and stabilized by two screws under three different conditions of cement augmentation: (1) group injected with CPC; (2) group injected with polymethylmethacrylate (PMMA) cement; and (3) control group without cement augmentation. The CPC and PMMA cement were injected into the screw holes before the screw fixation. At 3, 6, and 12 weeks after the screw fixation, the sheep was sacrificed and the proximal femur removed as the test specimen. Applying compressive force on the femoral head, the maximum load to fixation failure was measured using an Instron testing machine. Maximum load to fixation failure at 3 weeks after screw fixation was highest for the PMMA group compared with the CPC and control groups. At 12 weeks after screw fixation, the value was highest for the

Figure 4 X-ray diffraction analysis in vitro and in vivo shows similar pattern—12 peaks of -TCP, TTCP, and HA after 3 days. Peak of -TCP and TTCP decreased, and peaks of HA increased gradually after 7 days. Peaks of -TCP and TTCP mostly diminished, and peaks of HA developed for 4 weeks. The crystallization in vivo was delayed about 3 weeks compared with the crystallization in vitro. (Black arrow: hydroxyapatite; white arrow: -TCP.)

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Figure 5 The CPC shows carbonate apatite pattern by the FTIR spectrum on 7 days after mixing in vitro (a) and in vivo (b).

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(a)

(b) Figure 6 (a) Microscopic findings at 4 weeks after injection of the CPC. (Left) The newly formed bone attached to the CPC. (Right) The newly formed bone (yellow color) was observed in surrounding bone by tetracycline labeling method. (b) Microscopic findings at 4 weeks after injection of the CPC. (Upper left, right) HE staining, CMR. The mesh-shaped new bone was observed surrounding the CPC. There is no giant cell due to foreign body reaction and fibrous tissue between bone and the CPC. (Lower left) The newly formed bone invades into the CPD (Alcian-blue staining).

CPC group that showed stable and rigid fixation (Fig. 10). Histological observation showed good bone conduction and no gap between the bone tissue and the injected CPC. In contrast, the PMMA group showed thick fibrous tissue surrounding the screws. The control group without cement augmentation showed ordinary osteointegration (Fig. 11). 5 Clinical Trials From October 1995 to November 1997, clinical trials were performed by experienced orthopedic surgeons (who were previously instructed on the use of CPC) at subsidiary hospitals of Aichi Medical University, Osaka City University, Kitasato University, Kochi University, and Hamamatsu University School of Medicine. During the trial period, the CPC was applied to 59 female and 30 male patients. The clinical cases were benign bone tumor,

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fracture in aged people, bone defect after harvesting graft bone (as a filler), and fracture (as augmentation for fixation devices). Results of 72 cases treated using comparable protocols for the clinical trials were evaluated. Radiographic results showed new bone formation around the CPC 1 to 3 months after surgery. The interfaces between the injected CPC and bone became less distinguishable with time. Side effects which could cause clinical problems were not observed. Sixty-six of 72 applications (91.7%) were classified as “very useful” or “useful.” These results indicate that the CPC is highly safe and effective as bone filler [42]. 6 Clinical Cases Clinical application of CPC in several cases of benign tumors [42,43] demonstrated that with time, the CPC was absorbed and gradually replaced by new bone (Fig. 12). The change of volume of injected CPC in seven cases of benign bone tumor (Table 2) was observed ra-

(a)

(b) Figure 7 (a) Scanning electron microscopy analysis shows 3 days after injection of the CPC into the medullary cavity of the rabbit femur. Macrophage phagocytes piece of the CPC. (Black arrow: CPC.) (b) The periphery of the CPC was covered by many osteoblasts. Newly formed bone is well developed. (Black arrow: CPC.)

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Figure 8 Microscopic findings (HE staining): small blood vessels extend into the CPC injected into the medullary cavity of rabbit femur at 8 weeks (black arrow).

Figure 9 Mechanical compression strength of the injectable CPC we investigated shows about 80 MPa in vitro and in vivo by 1 week after mixing and then remains steady for 4 weeks.

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Figure 10 Biomechanical evaluation: at 12 weeks after the screw fixation, CPC group maximum failure load showed the highest value compared with the other groups, which showed stable and rigid fixation.

diographically for 12 to 24 months (Fig. 13). Results suggest that the absorption rate of CPC depended on the following factors: (1) age of the patient (young or old), (2) site of implantation or injection (humerus, femur, etc.), and (3) the type of disease (bone cyst, enchondroma, etc). In a case of first lumbar spine compression fracture of a 68-year-old female, CPC injection was made transpedicularly into vertebral body after applying corrective cast [43]. Radiographs (Fig. 14) showed an increase in vertebral body height from 64% (preoperatively) to 85% (postoperatively) and 81% (24 months after operation). A significant increase in bone density was also observed.

Figure 11 Histological observation (Fuchsin staining): CPC group showed no gap between bone tissue and the injected CPC and good bone conduction. In contrast, PMMA group showed thick fibrous tissue surrounding the screws. Control group without cement augmentation showed ordinary osteointegration.

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Figure 12 Radiograph shows bone cyst in humerus of 13-year-old boy. The CPC was injected into the central part of the cyst, and autologous bone was implanted into the proximal and distal parts of the cyst. The CPC was absorbed and gradually replaced by new bone.

B General Clinical Applications for Injectable Calcium Phosphate Cements Injectable calcium phosphate cement has provided a wide indication for clinical applications in filling bone defects [27–29,32,34–48]. These applications include (1) benign bone tumor—as filling material for bone defect after curettage of tumor tissue (bone cyst, enchondroma, benign giant cell tumor, etc.); (2) fracture—as augmentation material for fracture fixation, especially osteoporotic bone fracture (tibia plateau compression fracture, cal-

Table 2 Cases of Benign Bone Tumor Injected with CPC Case no.

Sex

Age (years)

Site

Tumor type

1 2 3 4 5 6 7

Male Male Male Female Female Female Female

13 40 14 22 15 33 29

Humerus Femur Radius Finger Finger Femur Femur

Bone cyst Bone cyst Bone cyst Enchondroma Enchondroma Aneurismal cyst Giant cell tumor

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Figure 13 Radiographic demonstration of the change of volume of injected CPC in seven cases of benign bone tumor, 12 to 24 months after injection. (From Ref. 44.)

caneus fracture, femoral neck fracture, distal end of radius fracture, thoracolumbar spine fracture, etc.; (3) substitution material for bone defect after harvesting of autologous bone graft in the iliac crest; and (4) substitution material for filling gaps between stem of uncemented prosthesis and bone. In addition to the clinical applications cited, CPC may be used as a delivery system for antibiotics, anticancer drugs, anti-inflammatory drugs, and as a carrier and delivery system for bone morphogenetic protein or biologically active peptides. It may also be a useful material as a scaffold for tissue engineering for the regeneration of bone tissues.

Figure 14 Radiograph of lumbar spine fracture shows the first lumbar spine compression fracture of 68 females. The CPC was injected transpedicularly into vertebral body after applying corrective cast. Vertebral body height was 64% preoperatively, 85% postoperatively, and 81% 24 months after operation. An increase in bone density was also noted.

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III CONCLUSION Injectable calcium phosphate cement has special characteristics as a biomaterial for bone repair: (1) it has composition, crystallinity, and bioresorbability similar to the bone mineral; (2) it has similar or higher mechanical strength than normal cancellous bone; and (3) it is osteoconductive and bioactive and allows its replacement by the new bone by a process similar to the remodeling of normal cortical and cancellous bone. ACKNOWLEDGMENTS The authors gratefully acknowledge the contributions of the Bio-research group, Department of Orthopedic Surgery; Dr. Norio Kawai, Department of Anatomy, Aichi Medical University; Professor Xin Xiang Su, Norman Bethune University of Medical Sciences, Changchun, China; as well as the contribution of Dr. Shuji Lin and Dr. John P. LeGeros, Calcium Phosphate Research Lab and Department of Biomaterials, New York University College of Dentistry. The injectable calcium phosphate cement used in the clinical studies cited in this chapter was donated by the Mitsubishi Material Co., Japan. REFERENCES 1. Albee F. H. 1920. Studies in bone growth. Triple calcium phosphate as stimulus to osteogenesis. Ann. Surg. 71:32–36. 2. Nery E. B., Lynch K. L., Hirthe W. M., Mueller K. H. 1975. Bioceramics implants in surgically produced infrabony defects. J. Periodontol 46:328–339. 3. LeGeros R. Z. 1988. Calcium phosphate materials in restorative dentistry: a review. Adv. Dent. Res. 2:164–183. 4. LeGeros R. Z., Daculsi G. 1990. In vivo transformation of biphasic calcium phosphate ceramics: ultrastructural and physico-chemical characterizations. In: Handbook of Bioactive Ceramics, Vol. II: Calcium Phosphate Ceramics, Yamamuro N., Hench L., Wilson Hench J., Eds. CRC Press: Boca Raton, FL, pp. 17–28. 5. Nery E., LeGeros R. Z., Lynch K. L. 1992. Tissue response to biphasic calcium phosphate ceramic with different ratios of HA/-TCP in periodontal osseous defects. J. Periodontol 63:729– 735. 6. Daculsi G., Bagot D’Arc M., Corlieu P., Gersdorff M. 1992. Macroporous biphasic calcium phosphates efficiency in mastoid cavity obliteration. Ann. Otol. Rhinol. Laryngol. 101:669– 674. 7. Passuti N., Daculsi G., Martin S., Deudon C. 1990. Macroporous calcium phosphate ceramics for long bone surgery in human and dogs. Adv. Biomater. 9:255–258. 8. Alam M. I., Asahina I., Ohmamiuda K., Enomoto S. 2001. Comparative study of biphasic calcium phosphate ceramics impregnated with 4hBMP-2 as bone substitutes. J. Biomed. Mater. Res. 54:129–138. 9. Jarcho M. 1981. Calcium phosphate ceramic as hard tissue prosthetics. Clin. Orthop. 157:259– 278. 10. Aoki H. 1991. Science and Medical Applications of Hydroxyapatite. Japan Association of Apatite Science (JAAS), Takaya Press System Center: Tokyo. 11. DeGroot K., Ed. 1983. Bioceramics of Calcium Phosphate. CRC Press: Boca Raton, FL. 12. deJong W. F. 1926. La substance mineral dans les os. Tec. Trav. Chim. 45:445–458. 13. Kay M. I., Young R. A., Posner A. 1964. Crystal structure of hydroxyapatite. Nature 294: 1050–1053. 14. LeGeros R. Z. 1981. Apatites in biological systems. Prog. Crystal Growth Charact. 4:1–45.

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35. Bai B., Jazrawi L. M., Kummer F. J., Spivak J. M. 1999. The use of an injectable, biodegradable calcium phosphate bone substitute for the prophylactic augmentation of osteoporotic vertebrae and the management of vertebral compression fractures. Spine 24(15):1521–1526. 36. Bohner M. 2000. Calcium orthophosphates in medicine: from ceramics to calcium phosphate cements. Injury Int. J. Care Injured 31:S-D37–47. 37. LeGeros R. Z. 1993. Biodegradation and bioresorption of calcium phosphate ceramics. Clin. Mat. 14:65–88. 38. Hirano M. 1993. Development of -TCP bioactive cement. New Ceramics [Japanese] 5:55–59. 39. Sugimoto T., Sato K., Morikawa K., Kawai N., Niwa S. 2000. Histological examination of the calcium phosphate with vascularization. J. Jpn. Orthop. Assoc. 74:S1657. 40. Yamamoto H. 1997. Experimental studies for clinical application of calcium phosphate cement. J. Aichi Med. Univ. Assoc. 25:219–229. 41. Zhang W. 2000. Basic research and clinical application of augmented screw fixation with calcium phosphate bone cement (CPBC) for the proximal femoral fractures. PhD thesis, Norman Bethune University of Medical Science, China. 42. Yamamoto H., Mixobuchi H., Sibata T., Niwa S., Oota H., Morikawa K., Yoshiki Y., Nichimura N., Hidaka N., Itoman M., Sekiguchi M., Inoue T., Miyamoto S. 1998. Clinical evaluation of calcium phosphate bone past (CPC 95) in orthopedic surgery. Jpn. Pharmacol. Ther. 26:409–429. 43. Yamamoto H., Shibata T., Ikeuti M. 1999. Calcium phosphate cement injection for osteoporotic vertebral fracture. Clin. Orthop. Surg. [Rinsho Seikei Geka—Japanese] 34:435–442. 44. Oota H., Sato K., Morikawa K., Sugimoto T., Niwa S. 1998. Calcium phosphate cement for curetted benign bone tumor. Cent. Jpn. J. Orthop. Traumatol. 41:1269–1270. 45. Morikawa K., Niwa S., Ohta H., Hanabayasi A., Sato K., Siohara H., Sumita S., Hirano M. 2000. The countermeasure for complication hip screw fixation: the application of calcium phosphate cement as an augmentation material for lag screw fixation. Seikei-geka [Japanese] 37(Special Issue):111–114. 46. Hidaka N., Yamano Y. 1998. Fracture treatment by calcium phosphate cement. Orthop. Surg. Traumatol. [Japanese] 41:1387–1392. 47. Moore D. C., Maitra R. S., Farjo L. A., Graziano G. P., Golstein S. A. 1997. Restoration of pedicle screw fixation with an in situ setting calcium phosphate cement. Spine 22:1696–1705. 48. Mermelstein L. E., Mclain R. F., Yerby S. A. 1998. Reinforcement of thoracolumber burst fractures with calcium phosphate cement—a biomechanical study. Spine 23:664–671.

21 Inorganic Bone Substitutes R. Schnettler Justus-Liebig-Universität, Giessen, Germany E. Dingeldein Osartis GmbH & Co., Obernburg, Germany

I INTRODUCTION Bone transplants and bone substitute materials are necessary in about 10% of all reconstructive operations of the locomotor system caused by traumatic, resectional, or congenital defects. Time-honored clinical experience has shown that extensive bone defects often occur in trauma patients, requiring bony reconstruction. In operative fracture treatment, the transplantation of bone in the form of spongiosa has found a broad spectrum of applications in filling up defects and bridging zones of destruction. Here the higher osteogenic potential of autografts in comparison to allogenic transplants is undisputed, but is restricted by the limited availability and the necessity of a second operation, so that especially in reconstruction of greater defect sections allogenic donor bone is frequently employed. Allogenic bone transplants not only demand access to a costly and existing bone bank that has to be run according to governmental regulation, the transplantation of allogenic material also carries the potential risk of transmission of bacterial or viral infections, including the transmission of HIV. Additionally, there is a possibility of inoculation of tumor cells as well as of development of antibodies within the ABO system following blood group–incompatible bone transplantation with the risk of occurrence of Morbus haemolyticus neonatorum in the case of a later pregnancy. Principally, a more or less distinct histoincompatibility due to transfer of immunocompetent lymphatic donor cells exists in allogenic bone transplantation as well as in the transplantation of solid organs [1]. Since the fundamental examinations of osteoinduction and the affiliated isolation of growth factors by Urist in 1965 [2], extensive scientific research on the growth factors 401

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contained in bone matrix has been performed. Proteins of the TGF- family, e.g., the bone morphogenetic proteins (BMPs), play a key role in the regulation of bone regeneration and are therefore of special interest [3–5]. In the past years, the alkaline fibroblast growth factor has also raised increased interest among a great many researchers. Its presence in bone matrix leads to the supposition that it plays an important role in the development and growth of bone substance [6,7]. One of the best known effects is the significant augmentation of microangiogenesis, which could be demonstrated among others in experimental investigations on wound healing [8]. Further experimental examinations showed a significant increase of callus formation in rats and miniature pigs in which FGF had been injected into the fracture site [9,10]. The currently available bone substitute materials are only to be used in clearly defined indications, as they do not currently meet the biological or mechanical properties of autogenous bone. Our knowledge is grounded on various experimental models which are not always comparable. Therefore many aspects have to be considered as a working understanding. This chapter should be read with this in mind, and its length precludes an exhaustive review. According to today’s knowledge, a bone substitute material should have the following qualities: good local and systemic compatibility, no allergenic potential, capability of substitution by bone and the complete filling of any defect, osteoconduction, and stimulation of bone healing, or osteoinduction. The bone that in the end substitutes the artificial material should be able to bear weight and, if possible, be lamellar bone. Besides the autogeneous bone presently available, commercial and experimental bone graft materials show a variety of compositions and properties, many of which are very different from those of bone. Among them inorganic calcium derivatives are frequently used. These materials include calcium phosphate ceramics [hydroxyapatite (HA): Endobon, Cerabone, Ceros 80; and tricalcium phosphate (-TCP): Ceros 82], calcium phosphate cements (-BSM, Biocement D), nanoparticular hydroxyapatite paste (Ostim), and calcium sulfate (Calcigen and Cavat). Depending on the nature of their interfaces with host bone, these materials are described either as bioinert or bioactive [11–14]. The physicochemical properties of these materials were compared using x-ray diffraction (XRD) and scanning (SEM) and transmission electron microscopy (TEM). Biological reactivity of the different materials was also compared in histological evaluations in animal models. Experimental and clinical studies to date have been encouraging, especially in metaphyseal defects. These and other studies are indicative of the potential utility of resorbable and nonresorbable inorganic materials as a bone graft substitute. II MATERIALS AND METHODS A Materials The materials used for the study included HA ceramics, HA cements, nonsintered HA, and calcium sulfate. They are listed in Table 1. These materials were investigated in different animal models. Surgically created bone defects using a diamond bone cutting system in miniature pigs, sheep, and rabbits were filled and histologically evaluated after different follow-up periods (see Table 2). The materials were characterised prior to implantation using chemical analyses, x-ray diffraction and scanning or transmission microscopy.

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Table 1 Origin and Composition of Inorganic Bone Substitutes Material Endobon (Biomet-Merck) Cerabone (Coripharm/AAP) Ceros 80 (Mathys) -BSM (Etex, 26 Biomet-Merck) Biocement D (Biomet-Merck) Ostim (Osartis/AAP) Calcigen (Biomet-Merck)

Origin

Composition

Porosity

Resorption

Bovine bone Bovine bone Synthetic

HA

Macroporous

Nonresorbable

HA

Macroporous

Nonresorbable

HA

Macroporous

Nonresorbable

Synthetic Synthetic Synthetic Synthetic

HA HA HA Calcium sulfate

In situ curing paste In situ curing paste Noncuring paste Micoporous

Resorbable Resorbable Resorbable Resorbable

The materials differed in surface morphology, as shown by SEM as well as by TEM. Figure 1 shows the interconnecting pore system of a bone derived hydroxyapatite ceramic. A close-up of the trabeculae demonstrates the size of the crystals and their tight connection (Fig. 2). Synthetic ceramics like Ceros 80 are similar in their crystal configuration, but differ in the size and number of pores. Blind ending pores are frequently found (see Fig. 20, 21). Calcium phosphate cements consist of particles of different size embedded in a cement matrix (see Fig. 23). In case of Biocement D irregular, long round-shaped incorporations are seen (see Fig. 29). Agglomerates of Ostim nanocrystals are shown in Fig. 3. Calcium sulfate needles after hardening are of variable size and configuration (Fig. 4). Beside the chemical analysis the x-ray diffraction is also an important method to determine the composition of different minerals because many chemical properties depend on the phases of the substances. With chemical analysis it is possible to determine atomic species, ions, or molecule groups but not the phases of the substances. X-ray diffraction has been in use in two main areas: for the fingerprint characterization of crystalline materials and the determination of their structure. Each crystalline solid has its unique characteristic x-ray powder pattern which may be used as a “fingerprint” for its identification. Once the material has been identified, x-ray crystallography may be used to determine its structure, the crystalline state, and the interatomic distance and angle. X-ray diffraction is one of the most important characterization tools used in solid state chemistry and materials science. Table 2 Inorganic Bone Substitutes and Animal Model Materials Endobon (Biomet-Merck) Endobon (Biomet-Merck) Ceros 80 (Mathys) -BSM (Etex/Biomet-Merck) Biocement D (Biomet-Merck) Ostim (Osartis/AAP) Ostim (Osartis/AAP) Calcigen (Biomet-Merck) Cavat (Coripharm/AAP) Cavat (Coripharm/AAP)

Species

Model

Follow-up

Miniature pigs Rabbits Miniature pigs Sheep Sheep Sheep Rabbit Sheep Sheep Rabbit

Femur condylus, press-fit Femur condylus, press-fit Femur condylus, press-fit Tibial head Tibial head Tibial head Femur condylus Tibial head Tibial head Femur condylus

42 days, 84 days 42 days, 84 days 42 days, 84 days 42 days 42 days 30 days, 60 days 20 days, 40 days 42 days 30 days 14 days, 60 days

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Figure 1 HA ceramic derived from bovine bone.

The size and the shape of the unit cell for any compound can be determined most easily using the diffraction of x-rays, and an unknown mineral can be identified or the atomic-scale structure of an already identified mineral can be characterized. Systematic x-ray diffraction data for thousands of mineral species exist. A large amount of these data has been collected and published by the JCPDS International Centre for Diffraction Data. The x-ray diffraction patterns of bone derived and synthetic hydroxyapatite are nearly identical (Figs. 5 and 6).

Figure 2 Higher magnification of HA crystals and their connections.

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Figure 3 Precipitated Ostim nanocrystals.

Figure 4 Calcium sulfate needles.

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Figure 5 X-ray diffraction analysis of bovine bone–derived HA ceramic.

Figure 6 X-ray diffraction analysis of synthetic ceramic. 406

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Diffraction Angle (degrees") Figure 7 X-ray diffraction analysis of -BSM, natural bone, and hydroxyapatite. -BSM was compared to natural bone and hydroxyapatite. Although similar to natural bone, in comparison to hydroxyapatite diffraction peaks are distinctly broader (Fig. 7). The x-ray diffraction pattern of Biocement D was compared to synthetic hydroxyapatite; it is shown in Fig. 8. The diffractogram of Ostim hydroxyapatite paste demonstrates wider peeks due to the nanosize of the particles; it was compared to a hydroxyapatite standard (Fig. 9). The XRD of calcium sulfate powder (CaSO4)  0.5H2O revealed pure Bassanite (Fig. 10). B Animal Models Cylindrical bone defects were created using a diamond coated core drill [diamond bone cutting system (DBCS)] with water irrigation. The size of the defects were 9.6 mm in diameter and 10 mm in depth in miniature pigs, 11 mm in diameter and 20 mm in depth in sheep, and 5.6 and 8 mm, respectively, in rabbits.

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Figure 8 X-ray diffraction analysis of Biocement D.

Figure 9 X-ray diffraction analysis of Ostim hydroxyapatite.

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Figure 10 X-ray diffraction analysis of calcium sulfate.

Specimens were left in vivo for different time periods (see Table 2). Specimens were fixed in Karnowsky solution before being embedded in Technovit resin. Each specimen was sectioned longitudinally or horizontally and sections were ground to a thickness of 20–30 m using the “exact method” [15]. The sections were then stained with Toluidine blue for qualitative and quantitative examinations. Unstained sections were prepared for the study of fluorochrome labels. Semithin sections (1 m) were stained with Richardson (1% methylene blue, 1% borax, 1% azure II). Ultrathin sections (80 nm) were counterstained with uranyl acetate and lead citrate (Reichert Ultrostainer, Leica, Germany) and examined in a Zeiss EM 109 transmission electron microscope. In order to study the nature of the new bone formation at defined time intervals fluorochrome labeling was employed during the loading period. The fluorochrome labels used were tetracycline, Alizarin complexon, Calcein green and Calcein blue. After sacrifice of the animals the bones carrying the implants were dissected by disarticulation of the knee. Ultrastructural analysis was performed on -BSM and Biocement D implants. C Results and Discussion The histological response to implants of different chemistry and structure was similar in ceramic HA implants but differed considerably when resorbable materials were im-

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Figure 11 Bovine bone–derived hydroxyapatite with areas of bone deposition.

planted. In sintered HA ceramics (Fig. 11) active areas of bone deposition, as well as resorption and remodeling, were present. By 5 weeks bone ingrowth started (Fig. 12) and was complete after 10 weeks when the bone around the fully integrated implant underwent normal remodeling with osteoclastic (Fig. 13) and osteoblastic (Fig. 14) activity. Abundant evidence of osteoblastic activity was seen in the early stages, but particularly

Figure 12 Hydroxyapatite ceramic in miniature pigs 5 weeks after implantation, starting bone ingrowth.

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Figure 13 Active area of bone resorption by osteoclasts.

at times where appositional activity was present. Bone ingrowth tended to proceed from the bottom of the defect as well as from the walls (Fig. 15). There were no signs of fibrous encapsulation, but in some sections aggregations of loose HA cristallites surrounded by macrophages were observed (Fig. 16). This is directly related to the implantation procedure, however due to their porosity, osteoconductive HA implants can undergo a significant degree of degradation and resorption. Fluorochrome labeling revealed that ingrowth/ongrowth of bone at the external surfaces of the implants which occurred after 4–6 weeks tended to be of lamellar nature with the clear apposition of bone

Figure 14 Formation of appositional new bone by osteoblasts.

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Figure 15 Bone ingrowth into the ceramic implant starting from host bone at the bottom and the walls.

directly on internal implant surfaces (Figs. 17 and 18). We noticed a transient appearance of multinucleated giant cells (Fig. 19). After 1 month vascularization of the bone within the implant occurred when an interconnecting pore system was present. Implants with closed pores showed little evidence of bone ingrowth (Fig. 20). In general the tissue response to porous HA implants is different from that to dense

Figure 16 Loose HA crystals at the ceramic surface surrounded by macrophages.

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Figure 17 Lamellar bone ongrowth at the external surface of HA implants.

HA. The amount and the type of bone ingrowth is dependent on the porosity and the interconnectivity (Fig. 21). The qualitative results described show that HA ceramic with an interconnective pore system is highly osteoconductive, acting as a framework for appositional bone ingrowth. When bone ingrowth is complete throughout the pores of an osteoconductive material, the implant–bone composite significantly changes the original biomechanial properties. There is a high correlation between the amount of bone ingrowth into the interconnective pores and the bending strength of porous HA [16,17].

Figure 18 Direct bone apposition at the ceramic surface.

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Figure 19 Multinucleated giant cells and HA crystals. Investigations into the basic characteristics of calcium phosphate cements revealed several material properties that appear important for clinical applications. Calcium phosphate cements are provided as a powder that is hydrated with water, buffer solution, or physiological saline to form an injectable paste which hardens at body temperature. The mixing initiates the setting reaction that transforms the powder into a substance of pastelike consistency. The paste is applied to the bone defect. This property is a major advance over ceramic granula HA that depends upon defect geometry and additional surgical techniques to contain the granules in the defect. Calcium phosphate cements adhere to bone and the inherent setting stabilizes the implant.

Figure 20 Synthetic HA ceramic (Ceros 80) with bone ingrowth.

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Figure 21 Synthetic HA ceramic (Ceros 80); no bone ingrowth into blind ending pores. Using a sheep tibia defect model, we could demonstrate degradation and resorption of calcium phosphate cements by osteoclasts preceding new bone formation by osteoblasts. The comparison of -BSM (Biobon) and Biocement D after 6 weeks showed a more pronounced reaction in the -BSM group (Figs. 22,23). The semithin section stained with Richardson (1% methylene blue, 1% borax, 1% azure II) shows long slender projections of bone marrow—including blood vessels and associated cells—forming a circumscribed labyrinthine system within the cement. The outer surface of the bone marrow facing the cement is covered by a cellular band of cuboidshaped osteoblasts, the nuclei of which generally contain two peripherally located nucleoli. Adjacent areas of collagen fibers arising from the osteoblasts are closely associated with the cement. The thickness of this compact collagen “sheath” varies from one region to an-

Figure 22 -BSM resorption channels.

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Figure 23 Higher magnification of -BSM resorption channels and bone apposition. other. In the upper left corner of Fig. 24 a lipocyte revealing a large lipid droplet is surrounded by two capillaries. The capillary of the bone marrow projection contains a monocyte; osteoclasts are not localized in this section plane. A transmission electron micrograph shows two adjacent osteoblasts responsible for the synthesis of the irregularly arranged bundles of collagen fibers. These are closely associated with the ceramic surface and also occupy some of the small spaces within the calcium phosphate cement. The ultrastructural features of the osteoblasts reveal that the nucleus is ovoid or ellipsoidal and in general eccentrically located. The chromatin is homogenously dispersed throughout the karyoplasm, except for slightly condensed clumps which are arranged along the inner aspect of the nuclear membrane. Within the cytoplasm the rough endoplasmatic reticulum is the most prominent component, with the cisternae arranged in parallel arrays.

Figure 24 -BSM semithin section with cuboid osteoblasts and osteoid.

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Figure 25 -BSM; ultrastructure of osteoblasts. Elongated mitochondria with dense internal matrices and well-developed Golgi complexes are also visible (Fig. 25). Higher magnification demonstrates long profiles of the rough endoplasmic reticulum, extending throughout the cytoplasm of the apparent osteoblast. Golgi complexes are composed of multiple stacks of flattened, smooth-surfaced cisternae. Mitochondria and a single centriole can also be seen. In the upper half of the micrograph, irregularly arranged collagen fibers are visible. The bundles are in close contact with the ceramic surface (Fig. 26). The multinucleated osteoclast closely associated to the ceramic surface is cuboidal in shape and reveals clearly defined plasma membrane domains such as the ruffled border, the sealing zone, and the dorsal microvilli. Ultrastructural features demonstrate that the osteoclast displays intended nuclei which are located toward the base of the cell. Dense clumps of chromatin are prominent within the karyoplasm and occur in a small peripheral band. The cytoplasm is occupied by numerous mitochondria and membrane-bound vacuoles differing in shape, size, and content. Large vacuoles filled with long slender crystals are preferentially localized in vicinity to the ruffled border, whereas smaller vacuoles enclosing spherical particles are located below. These vacuoles could be involved in selective uptake of cement material from the extracellular medium and point toward a degradation mechanism of osteoclasts by simultaneous resorption and phagocytosis (Fig. 27). This degradation mechanism has been described in a recent in vitro study [18]. Ultrastructural details of the osteoclastic ruffled border facing the ceramic surface show mitochondria and vacuoles which lie adjacent to the ruffled border. Long, slender crystals are in close contact with the cellular membrane of the ruffled border and can also be seen within a larger vacuole whose membrane is in continuity with the ruffled border. Small and rounded crystals are encircled by smaller vacuoles. In the upper quarter of Fig. 28 the calcium phosphate cement can be identified due to its crystalline appearance and high electron density. Note the alterations of shape and electron density of cement crystals which are located close to the ruffled border, indicative of the intense resorption activity of the cell. These changes in size and shape of the crystals might be due to in situ fragmentation. In case of Biocement D, as far as strength is concerned for load-bearing applications the hardened cement is in the same range as cancellous bone when compressive strength is compared [19].

Figure 26 -BSM; higher magnification of ultrastructure of osteoblast.

Figure 27 -BSM; ultrastructure of osteoclast. 418

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Figure 28 -BSM; higher magnification of ultrastructure of osteoclast. Histological results obtained in sheep confirm biocompatibility and bone ongrowth to the implant after 6 weeks. A difference in density shows the beginning of implant resorption (Fig. 29). In a few areas partial resorption of the implant and direct osseointegration can be observed (Fig. 30). Semithin microtome section illustrated the cellular populations involved in degradation of the cement and in the synthesis of the collagen fiber bundles. The bone marrow areas facing the cement surface are covered by multinucleated osteoclasts and by osteoblasts. The former are visible at the left margin of the tissue. The osteoclastic ruffled borders appear as lightly stained cellular regions which are closely attached to the cement surface. Numerous osteoblasts can be seen at the right margin of the bone marrow projection. They are in close apposition to each other and usually have roundish or oval profiles. The sectioned bone marrow is well vascularized, and in the vicinity of the endothelial tubes surrounded by cells of the osteoblastic lineage, a wide range of shapes and cytoplasmic densities due to the different maturational stages within this cell population are observed. A multinucleated giant cell localized in the upper left corner of the bone marrow projection can also be identified. Note the expanded area of host bone in the right half of the picture which is in close contact with the neighboring osteoblastic layer (Fig. 31). Higher magnification shows the intimate contact of a bone marrow column with the calcium phosphate cement and details of the cell populations involved in dynamic resorption and bone formation. Multinucleated osteoclasts are situated at the bone marrow– cement interface. Their ovoid-shaped nuclei contain dense clumps of heterochromatin

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Figure 29 Biocement D; beginning of implant resorption. which are irregularly distributed throughout the pale karyoplasm. The ruffled border is represented by the slightly stained cytoplasmic portions oriented in the direction of the cement. The remaining cell population of the well-vascularized bone marrow is composed of hemopoietic cells, osteoprogenitor cells, and fibroblasts, a representative part of which is localized in the lower half of the sectioned bone marrow. Note the multinucleated giant cell and the macrophage which usually occur scattered individually (Fig. 32). The mineral component of bone is predominantly hydroxyapatite, so the biocompatibility of synthetic hydroxyapatite is predictably good. The results depend upon the particle size and phase purity of the material. In case of Ostim paste the stochiometry of the product is of importance. The calcium phosphate ratio is 1.67 in pure hydroxyapatite and Ostim shows the same calcium phosphate ratio. Due to the nanosize crystals the surface area is very large, 100 m2/g, and thus very close to the surface area of bone mineral. The

Figure 30 Biocement D; resorption channel.

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Figure 31 Biocement D; semithin section. biomaterial Ostim is an injectable paste containing hydroxyapatite nanocrystals and water. It binds to bone tissue and stimulates bone healing in critical-size defect in animal models. In a rabbit tibia femur condylus model, Ostim promoted bone healing by 4 weeks. Since Ostim is not hardening in situ, cell infiltration can start immediately after revasculerization of the implant site. Fluorescent dyes incorporated in calcifiying tissues provide time marks in the study of bone growth, turnover, and repair of bone. When these dyes are present in the bloodstream they will deposit in areas of the skeleton where new bone matrix, both fibrous and lamellar, is mineralizing. Fluorochrome labels applied from day 15 and 18 were already present in area of implantation indicating that mineralization had started in the newly formed osteoid. The nanoparticulate hydroxyapatite paste Ostim was well tolerated in

Figure 32 Biocement D; semithin section at higher magnification.

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Figure 33 Ostim globuli surrounded by newly formed bone.

sheep and rabbits. Four weeks after implantation the Ostim paste was disintegrating into globuli and infiltrated by macrophages and osteoid-producing cells which covered the Ostim globuli partially (Fig. 33). Ostim paste was implanted into tibial head bone defects in sheep. New bone formation spread from the host bone into the defect. Around the implants an increase of trabecular bone could be observed (Figs. 34 and 35). After 4–6 weeks the defect was completely filled with new bone growing between and around the Ostim globuli. In cancellous bone, the filling of the defect became even more dense and compact after 12 weeks.

Figure 34 Ostim defect bridging after 4 weeks.

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Figure 35 Ostim defect bridging at higher magnification. Calcium sulfate has been used to fill bone defects for more than 100 years [20]. It is considered an inexpensive and safe material where absorption is followed by reconstitution of bone. In our animal studies in sheep the absorption of the calcium sulfate started immediately after implantation revealing an outer zone of disintegration of the implant followed by a zone of inflammation with areas filled with secretion and debris. The zone of inflammation was highly vascularized. New bone formation was only seen at a distance to the implant, spreading from the host bone toward the defect area. After 6 weeks signs of inflammation disappeared. The implant degraded from the outer surface. New bone formation growing from the host bone remained at a distance to the implant (Figs. 36 and 37).

Figure 36 Calcium sulfate; signs of inflammation and new bone formation.

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Figure 37 Calcium sulfate resorption after 6 weeks. In rabbit femur condylus, defects were completely filled with newly formed bone within 6 weeks, whereas in sheep after 12 weeks the center of the defect was still empty. Volumetric defects of bone are frequently encountered in the management of orthopedic trauma. While it is commonplace to implement autogenous bone grafting to fill these defects, this practice is not without the risk of added morbidity to the patient. The use of bone graft substitutes fills bony voids while reducing the morbidity encountered with autogenous bone grafting. These materials pose no risk of disease transmission as exist for the use of allograft bone. In clinical practice it is generally observed that different defect locations are associated with different levels of difficulty in forming new bone. Therefore in order to select the appropriate material out of a number of commercially available bone substitutes, their morphology, chemistry, and histological behavior in bone defect must be known in detail. Comparative studies can be helpful, but finally only validation of the bone graft subtitute in the pertinent clinical sites may be predictive of performance. D Clinical Cases Two representative clinical cases that needed bone grafting due to volume loss of bone at the fracture size are presented. 1 Case 1 A 72-year-old female was struck by a car and sustained a fracture of her right elbow joint. The fracture encountered a pre-existing bone cyst. She underwent open reduction, curettage of the cyst, and internal fixation of this injury. To avoid the harvest of autograft from the iliac crest, an HA ceramic was chosen to fill the bone cyst. After 8 months the radiograph exhibits evidence of fracture healing with incorporation of the HA ceramic (Endobon). A biopsy cylinder was taken with consent of the patient and processed histologically. Complete bone ingrowth had occurred with direct bone bonding to the ceramic trabecule (Figs. 38–44). (text continues on p. 428)

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Figure 38 Patient 1: fracture of the right elbow joint.

Figure 39 Patient 1: fracture of the right elbow joint (intraop situs).

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Figure 40 Patient 1: fracture of the right elbow joint received HA ceramic implant.

Figure 41 Patient 1: fracture of the right elbow joint received HA ceramic implant (postop x-ray).

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Figure 42 Patient 1: fracture of the right elbow joint received HA ceramic implant after removal of hardware and biopsy.

Figure 43 Patient 1: biopsy.

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Figure 44 Patient 1: histological section. 2 Case 2 A 48-year-old male was involved in a car accident and sustained multiple injuries including a comminuted, intra-articular, depressed calcaneous fracture of the right foot. Calcaneal fractures have to be addressed operatively with open reduction reconstruction of the depressed subtalar joint and internal fixation. It has been demonstrated in long-term studies that improvement in maintaining the subtalar joint space has decreased postoperative arthrosis. The depression was elevated via lateral approach. An associated large bone void was filled with -BSM and the reconstruction was held in position with a calcaneal-specific plate and a K wire. Fracture healing occurred after 6 weeks, and the metal was removed after 8 months. The calcium phosphate cement was radiographically visible at this time (Figs. 45–50). (text continues on p. 431)

Figure 45 Patient 2: calcaneal fracture CT-scan.

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Figure 46 Patient 2: calcaneal fracture bone void.

Figure 47 Patient 2: calcaneal fracture received -BSM implant.

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Figure 48 Patient 2: calcaneal fracture postop x-ray.

Figure 49 Patient 2: calcaneal fracture, postop CT-scan after hardware removal.

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Figure 50 Patient 2: calcaneal fracture received -BSM implant.

ACKNOWLEDGMENTS We would like to thank Mrs. S. Wenisch, PhD; Mrs. A. Hild, TA; and Mr. J. P. Stahl for their help, and companies Biomet-Merck and Coripharm for their financial support of the animal experiments. REFERENCES 1. Schnettler R., Dingeldein E., Tausch W., Ritter B. 1994. Untersuchungen zur knöchernen Integration einer Hydroxylapatit-Keramik und “bFGF” im Vergleich zu autogenen Spongiosazylindern. Osteo. Int. 2:118–126. 2. Urist M. R., Silverman B. F., Büring K., Dubuc F. L., Rosenberg I. M. 1967. The bone induction principle. Clin. Orthop. 53:243. 3. Urist M. R. 1968. Surface-decalcified allogenic bone (SDAB) implants. Clin. Orthop. 56:37. 4. Reddi A. H. 1981. Cell biology and biochemistry of enchondral bone development. Collagen Res. 1:209–226. 5. Reddi A. H., Wientroub S., Mutukumaran N. 1987. Biologic principles of bone induction. Orthop. Clin. N. Am. 18:207–213. 6. Globus R. K., Petterson-Buckendahl P., Gospodarowicz D. 1988. Regulation of bovine bone cell proliferation by fibroblast growth factor and transforming growth factors. Endocrinology 123:98–105. 7. Gospodarowicz D., Greenburg G. 1979. The effects of epidermal and fibroblast growth factors on the repair of corneal endothelial wounds in bovine corneas maintained in organ culture. Exp. Eye Res. 28:147–157. 8. Massagué I. 1987. The TGF family of growth and differentiation factors. Cell 49:437–438. 9. Jingushi S., Heydemann A., Kana S. K., Macey L. R., Bolander M. E. 1990. Acidic fibroblast

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Schnettler and Dingeldein growth factor injections stimulates cartilage enlargement and inhibits cartilage gene expressions in rat fracture. J. Orthop. Res. 8:364–371. Aspenberg P., Wang J. S. 1994. Basic fibroblast growth factor. Dose and time dependance in rats. Trans. Orthop. Res. Soc. 19:181. Revell P. A., Hing K. A., Tanner K. E., Best S. M., Bonfield W. 1997. Osteointegration of porous hydroxylapatite. In: Knochenersatzmaterialien und Wachstumsfaktoren, Schnettler R., Markgraf E., Eds. Thieme Verlag: Stuttgart, pp. 28–30. Daculi G., Bouler J. M. 1997. Bone bioactive ceramic interface—a dynamic process. In: Knochenersatzmaterialien und Wachstumsfaktoren, Schnettler R., Markgraf E., Eds. Thieme Verlag: Stuttgart, pp. 23–27. LeGeros R. Z., LeGeros J. P. 1997. Bone substitute materials and their properties. In: Knochenersatzmaterialien und Wachstumsfaktoren, Schnettler R., Markgraf E., eds. Thieme Verlag: Stuttgart, pp. 12–18. Osborn J. F. 1979. Biowerkstoffe und ihre Anwendung bei Implantaten. Schw. Mschr. Zahnheilkd. 89:1138–1139. Donath K. 1988. Die Trenn-Dünnschliff-Technik zur Herstellung histologischer Präparate von nicht schneidbaren Geweben und Materialien. Präparator 34:197–206. Martin R. B., Chapman M. W., Holmes R. E. 1989. Effects of bone ingrowth on the strongth and non-invasive asessment of a coralline hydroxylapatite material. Biomaterials 10:481–488. Hing K. A., Best S. M., Tanner K. E., Bonfield W. 1997. Materials in Medicine. J. Mater. Sci. 8:731–736. Heymann P. et al. 2001. Ultrastructural evidence in vitro of osteoclast-induced degration of calcium phosphate ceramic by simultaneous resorption and phagocytosis mechanisms. Histol. Histopathol. 16:37–44. Driessens F. C. M., Boltong M. G., De Maeyer E. A. P., Verbeeck R. M. H., Wenz R. 1998. Effect of temperature and immersion on the setting of some calcium phosphate cements. J. Mater. Sci. Med. (in press). Dreesmann H. 1892. Ueber Knochenplombierung. Beitr. Klin. Chir. 9:804–810.

22 Demineralization and Perforation of Cortical Bone Allografts: Preparatory Methods Kai-Uwe Lewandrowski, William W. Tomford, and Henry J. Mankin Massachusetts General Hospital, Boston, Massachusetts

I INTRODUCTION Large cortical autografts and allografts have been useful in limb-sparing procedures. They have frequently been applied to the treatment of bone tumors and for bone grafting in failed joint arthroplasties. These types of bone grafts, however, have been found to incorporate slowly into host bone resulting in susceptibility to nonunion, fatigue fracture, and infection [1–7]. Methods of enhancing incorporation of these types of grafts into host bone have included demineralization and perforation. Although deminerlization has been successful in improving osteoinductive properties of bone, as shown by the abundance of pertinent literature on experimental [8–11] and clinical studies [12–20], fully demineralized cortical bone has found limited clinical application. This appears primarily attributable to the fact that cortical bone, once subjected to extensive demineralization, loses its essential biomechanical properties. Sufficient mechanical strength is necessary in order to withstand skeletal forces when used in long bone reconstruction. In addition, demineralization of cortical bone yields a geometrical surface configuration that is less advantageous for bony ingrowth compared with cancellous bone, because cortical bone is less porous and has a comparatively low surface area–to–volume ratio. Recent studies have demonstrated the use of a mechanical drill [21] or an erbium:yttrium-scandium-gallium-garnet (Er:YSGG) laser [22–24] in cortical bone allografts, thereby increasing the porosity and allowing demineralization to proceed to areas that 433

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would normally be inaccessible to the demineralization process. When reimplanted, these grafts may therefore be more osteogenic than cortical grafts without holes. In an attempt to improve osteoinduction in cortical bone grafts, a new experimental approach using the erbium:yttrium-aluminum-garnet (Er:YAG) laser and partial demineralization was developed [22–25]. The pulsed mid-infrared Er:YAG laser was chosen because of minimal thermal damage (10–15 m) imparted to ablated bone [24,25]. In addition, it allows a large number of very small holes to be drilled into the grafts of this experimental model. Partial demineralization was employed to allow use of the grafts in weight-bearing implantation sites such as in long bones. In this chapter, the methods of partial demineralization and laser perforation of cortical diaphyseal bone allografts and how these processing methods affect their mechanical properties are described. Moreover, a detailed description of how the mechanical properties of these types of grafts can be predicted as a function of demineralization depth and extent of perforation has been provided. II THE REACTION FRONT IN ACID DEMINERALIZATION OF CORTICAL BONE Although demineralization of cortical bone is extensively used in the processing of human cadaver bones to produce osteogenic material and in the preparation of bone specimens for histological study, the microscopic morphology of the demineralization process has not been described. Several investigators have characterized the kinetics of the demineralization process of cortical bone commonly used in the form of grafts in dental [26,27] and orthopedic applications [28], and in the form of ground bone for the production of gelatin in industrial applications [29]. Others have described demineralization kinetics in the context of tissue processing for histology [30,31]. However, all of these studies have used the assumption of the “classic shrinking-core reaction model” for their mathematical analyses and have not based their investigations on actual microscopic observations. The theory of the classic shrinking-core reaction model has been generally applied for modeling purposes of fluid–solid diffusion systems [31] and implies the existence of a sharp reaction front that, in the case of demineralization of cortical bone, separates the demineralized from the mineralized portion. The reaction front is assumed to advance with increasing immersion time in the decalcifying agent, which implies the kinetics of this reaction front can be predicted for some ideal geometries by simple mathematical models [29]. In turn, this would allow implementation of the concept of controlled demineralization, which may have applicability to bone grafts in current clinical use [28]. The microscopic geometry of the demineralization process has been studied using scanning electron microscopy. The sharpness of the interface between mineralized and demineralized bone was demonstrated. The ultimate objective of this study was to verify the classic shrinking-core reaction model to diaphyseal cortical bone specimens. A Bone Sample Preparation The experimental bones used in this study were mature Sprague-Dawley rat tibias provided by Taconic Laboratories (Germantown, NY). These were stripped of soft tissues including the periostium and cut such that a 5-mm-long mid-diaphyseal portion was obtained. The bone marrow was removed by multiple washings in saline. Before demineralization, the ends of the rat tibia diaphyses were capped with acidresistant paraffin so that demineralization could only proceed from the periosteal to the en-

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dosteal surface. This was done in order to have the sample approximate the ideal “cylindrical geometry” in which no demineralization occurs at the ends of the cylinder. Demineralization was then performed in a 0.5 N hydrochloric acid bath at volume excess for various immersion times. The acid bath was kept on a heat plate at 25°C and stirred throughout the demineralization process. The temperature of the acid bath was monitored and held constant. A total of 20 diaphyseal bone specimens divided in four groups of five specimens each was included in this study. Bone specimens of group 1 were removed from the acid bath after 10 min, and specimens of group 2 and 3 after 20 and 60 min, respectively. Bone specimens of group 4 served as controls and were not subject to demineralization. To stop the demineralization process, bone specimens were then placed in a well-stirred volume excess distilled water bath for 6 h, with the water being changed after each 2 h. After partial demineralization, specimens were processed by fixation overnight in 4% glutaraldehyde in 0.1M cacodylate buffer, then rinsed in buffer, dehydrated in graded alcohol, infiltrated with propylene oxide and epon and embedded in epon. For scanning electron microscopy, approximately 1-mm thick cross-sections of the diaphyseal bone specimens were cut with a jeweler’s saw, mounted on a carbon planchet, and secured to a stub. Specimens were carbon coated in an Edward’s carbon evaporator. The SEM was performed using an AMRAY 1400 scanning electron microscope with a Kevex™ series 82 xray detector. Best visualization and photographic results were obtained using the backscatter detector. Identification of the calcium peak was obtained from an analysis of the x-ray spectrum. Generation of the dot maps was performed by creating a calcium “window” from the full spectrum. Control specimens of group 4 having no demineralization were processed in similar fashion. Electron micrograph (EM) scans of control bone specimens were obtained and used as a reference for comparison with the demineralized cortical bone and the background staining of the epon embedding. Photographic prints of EM-scanned specimens were used to determine the depth of the demineralization process measured as the distance from the periosteal surface to the position of the interface between the decalcified and the mineralized phase. The measurements were done around the circumference of diaphyseal bone specimens at angular increments of 30°. B Scanning Electron Microscopy On scanning electron microscopy of nondemineralized control bone specimens of group 4, the cross-sectional cylindrical geometry of the diaphysis was readily identified and found to be distinguishable from the background staining of the epon embedding. Scanning electron microscopy studies on partially demineralized bone specimens of groups 1, 2, and 3 revealed that in cylindrical geometry of diaphyseal bones the demineralization process takes the cross-sectional form of a circumferential band of uniform thickness that surrounds the inner mineralized cortical bone core (Fig. 1). This was confirmed by measuring the thickness of the demineralized phase at 12 different angular positions. The resulting demineralization depths presented in Table 1 are means of these measurements. During demineralization, the reaction front was noted to start at the outer periosteal surface of the cortex and to penetrate into the mineralized bone matrix. The interface between the demineralized and mineralized portion of the bone specimen was found to take the form of an advancing reaction front, as shown in Fig. 2a,c,e. The position of the reac-

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Figure 1 Scanning electron micrograph of a cross-section of a partially demineralized rat tibia diaphysis embedded in epon (E) showing a circumferential band of demineralized bone surrounding the inner undecalcified core. (* Indicates undemineralized cortical bone, bar 500 m.)

tion front was confirmed by evaluation of the corresponding calcium dot maps (Fig. 2b,d,f). The interface between the decalcified and the mineralized phase was found to be no thicker than 20 m and therefore to be extremely sharp. It can be considered a reaction front of bone mineral breakdown that is typical of a diffusion rate limited process. The process of acid demineralization of cortical bone was examined from a morphological standpoint in order to determine the sharpness of the reaction front believed to result from the demineralization process [28,29]. Rat tibia diaphysis was chosen as the investigated model because they were representative of the cylindrical geometry of long bones frequently used in current clinical practice, yet were small enough to be suitable for a high resolution analytic method such as scanning electron microscopy. Scanning electron microscopy has shown that hydrochloric acid demineralization of diaphyseal cortical bone specimens can be described by the shrinking core theory used for diffusion rate limited processes in fluid–solid systems [31]. The shrinkage of the inner unreacted cortical bone core is dependent on the immersion time in hydrochloric acid. On the basis of this study, it has been concluded that demineralization of cortical bone results in

Table 1 Demineralization Time and Depth

Figure no. Fig. 1 Fig. 2 Fig. 3a,b Fig. 3c,d Fig. 3e,f

Immersion time in HCl (min) Control, no demineralization 120 10 20 60

Mean measured demineralization depth (m) 0 264  32 71  14 94  18 153  26

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Figure 2 Scanning electron micrographs of cross-sections of rat tibia diaphyses (a,c,e) with corresponding calcium x-ray dot maps (b,d,f) demonstrating the reduction of the bone mineral containing cortical bone core (*). Short arrowheads point to the periosteal surface of the cortex; long arrowheads point to the interface. Demineralization was performed in 0.5 N HCl for increasing immersion time: (a,b) 10 min, demineralization depth  70 m; (c,d) 20 min, demineralization depth  95 m; (e,f) 60 min, demineralization depth  150 m. (Bar  100 m.)

an extremely sharp advancing reaction front separating the demineralized bone matrix from the inner cortical bone core. The interface at the reaction front was measured to be no thicker than 20 m, implying that the demineralization process of cortical bone is diffusion rate limited. The scanning electron microscopic studies on the microscopic morphology of cortical bone demineralization presented has supported the analysis of the demineralization kinetics by a diffusion model presented herein. This consequently allowed for the prediction of cortical bone demineralization provided that the diffusion equation given by Fick’s law could be solved for the geometry of a particular bone and that the diffusivity of acid into the demineralized bone is known [28–31]. These observations aided in the implementation of mathematical models to control the demineralization of cortical bone grafts. III KINETICS OF CORTICAL BONE DEMINERALIZATION The morphology and kinetics of hydrochloric acid demineralization of human and rat cortical bone were investigated with the objective of developing a method of controlled demineralization for structural bone allografts. Prior SEM studies have shown that the demineralization of cortical bone is a diffusion rate limited process with a sharp advancing reaction front. The demineralization kinetics of human cortical bone was determined as the

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advance of the reaction front versus immersion time by measuring extraction of bone mineral in both planar and cylindrical geometries. Partial demineralization of cortical bone requires the ability to precisely control the demineralization process, a method that has been lacking in previous studies [28]. A number of authors have approached the problem of controlling cortical bone demineralization, but no models directly applicable to the specific geometrical and tissue characteristics of human cortical bone allografts have been developed to date [4,18,21,32]. Therefore, for this investigation the morphology and kinetics of hydrochloric acid demineralization of cortical bone were systematically studied with the objective of producing a graft that may be rapidly incorporated yet allow stable reconstruction. Mathematical models based on diffusional mass transfer were developed to predict this process. A Demineralization Kinetics Studies The experimental bones used in this study were human and rat tibias. Human tibias from young healthy males were obtained from a bone bank that had been forced to discard the bones due to bacterial contamination. Mature Sprague-Dawley rat tibias were obtained from Taconic Laboratories. To investigate the demineralization kinetics of cortical bone, diaphyses of human tibias were cut into uniform discs of 5 mm diameter and 2 mm thickness. These discs were used to create two geometrical models by coating various surfaces of the discs with an acidresistant epon layer. A planar geometry model was created by coating the circumference of a disc (Fig. 3a), and a cylindrical geometry model was obtained by coating the top and bottom surface of a disc (Fig. 3b). This allowed the demineralization process to occur either from the top and bottom surface (planar geometry) or from the circumference (cylindrical geometry). Demineralization of these discs was performed by immersion in hydrochloric acid at concentrations of 0.5 N (1.875%w/v), 1 N (3.65%w/v), and 2 N (7.3%w/v) for various periods of time. An excess of acid volume was employed to avoid depletion of the acid concentration in the bath, and the acid bath was agitated to provide a uniform concentration. Demineralization of the bone discs was terminated by multiple washings in 4°C distilled water until pH 7 was reached. Bone discs were then dehydrated by immersion in 98% ethyl alcohol, anhydrous ether extraction, and air drying, 1 h for each unit operation, respectively. Using a Mettler AE 163 precision balance, the time course for bone mineral extraction immersion time in HCl was determined by measuring the dry weight of the discs before and after demineralization at different immersion times. Demineralization experiments were performed in triplicate. The normalized weight loss of the bone discs [see Eqs. (1) and (2)] was used to quantify the kinetics of the demineralization process. B Mathematical Modeling For modeling purposes, the kinetics of the demineralization experiments were analyzed using the diffusion equation with an advancing reaction front in both planar and cylindrical geometry. The analysis served to test the diffusion formalism to accurately predict the demineralization process of human cortical bone. As previously described [29,31,33], the assumptions of this model were (1) The demineralization of cortical bone with hydrochloric acid is a diffusion process in a fluid–solid system with a sharp advancing reaction front and can be described with Fick’s law. (2) The acid concentration gradient between undemineralized bone and bath is the driving force for

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Figure 3 (a) Planar model. Peripheral rim of the bone disc is epon coated. Demineralization proceeds from top and bottom surfaces. x 0  distance from outer surface to the center line; x  any given position between centerline and outer surface, i.e., the distance from the centerline to the reaction front;   penetration depth. (b) Cylindrical model. Top and bottom surfaces of the bone disc are epon coated. Demineralization proceeds from peripheral rim. r 0  the initial radius; r  any radial position; i.e., the radial position of the reaction front; rd  radius of the remaining inner undemineralized bone core;   penetration depth (  r0  rd).

the mass flux of acid into bone. (3) The acid concentration at the exposed outer surfaces of the bone specimens equals the bath concentration. (4) Acid diffuses into bone and reacts instantaneously with the bone mineral upon arrival at the reaction front. Thus, the process was considered to be diffusion rate limited and the concentration profile of acid within the demineralized bone matrix was treated as quasistatic.

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C Modeling for Planar Geometries To model the demineralization kinetics in planar geometry (Fig. 3a), the top and bottom surface of the bone discs have been considered as semi-infinite sheets. Consistent with the theoretical assumption of a quasi–steady state process, Fick’s law was used in a one-dimensional cartesian coordinate system: 2 m  0 x 2

(1)

where m is the mass concentration of hydrochloric acid (mass of acid/mass of solution) and x the axial position in the sample measured from the center line (Fig. 3a). The boundary conditions are given with an acid concentration of zero at the interface between demineralized bone matrix and cortical bone, m(x  x0  )  0 and an acid concentration at the outer surface of the bone disc equal to the bath concentration, m(x  x0)  m0. Integrating Eq. (1) and solving for the boundary conditions yields





x m(x,t)  m0 1   (t)

(2)

where m0 is the mass concentration of HCl in the bath, m is the mass concentration of HCl at any given position x in the sample, and (t) the depth of demineralization at time t. To determine the penetration depth (t) mass conservation at the reaction front x   is enforced as follows:  m M   D t x



(3)

x

where M is the molecular weight of HCl (g/mol), the moles of acid necessary to demineralize 1 cm3 of cortical bone (mol/cm3), D the mass diffusivity of HCl into the bone (cm2/s) and the density of the acid solution (g/cm3). Equation 3 relates the mass flux of acid into cortical bone with the rate at which the demineralization front proceeds. Taking the derivative of equation 2 and substituting into equation 3 yields upon integration (t) 

2 Dm0t

 M  .

(4)

The acid bath concentration C0 in mol/L is related to the mass concentration of HCl m0 as

m0103 C0   M

(5)

Substituting C0 for the mass concentration of the acid bath m0 yields the relation expressed in Eq. (3). Equation (5) may be expressed as   kt

(6)

where k  2 Dm0103/M is the slope of the linear function (t ). Using the value of k defined by the data and determined by titration, the effective mass diffusivity D of HCl into bone can be calculated with k 2 D   2C0

(7)

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D Modeling For Cylindrical Geometries In cylindrical coordinates, Fick’s law is expressed as 1  m   r   0 r r r





(8)

where r is the radius of a cylindrical bone. The penetration depth as a function of immersion time is dependent on the initial radius, r0. Therefore, it is advantageous to normalize for r0 and to perform the analysis with dimensionless variables. Equation (8) may be rewritten as 1  m*   r*   0 r * r* r*





(9)

where m*(r*)  m(r)/m0, r*  rd/r0, r*  rd/r0. d The boundary conditions become m*(r*  1)  1 and m*(r*  r*d )  0. Integrating Fick’s law for cylindrical geometry [Eq. (9)] and applying the boundary conditions yields



lnr * m*(r*)  1   lnr*d

(10)

To obtain a relation for the penetration depth of the demineralization process  as a function of the immersion time t, mass conservation at the reaction front is enforced giving rd m M    D  r t



(11)

rrd

Rewriting Eq. (10) using the dimensionless variables m*  m/m0, r*  rd/r0, t*  Dt/r 20 and   m0/M , a parameter related to the demineralization strength of the acid employed, yields r* m* d    t* r*



(12)

r*r* d

Taking the derivative of m* in Eq. (10) and substituting into Eq. (12) yields upon integration 2 2r*d 2 lnr* d  r* d 1  4t*

(13)

Equation (13) gives the radial position of the demineralization front rd in cylindrical geometry measured from the center line. In order to obtain a relation that determines the immersion time t necessary to demineralize to a certain penetration depth  with (t)  r0  rd(t), Eq. (13) can be rewritten by substituting C0 for the mass concentration of the acid bath m0 with Eq. (5) and solved for t as follows: W (t)  W   W W

(t)  r0 1 

0

(14)

where r0 is the initial radius. For both geometrical models, the data were normalized by the initial weight of the bone discs and presented as single data points. Figures 4 and 5 show the respective kinetic data for planar and cylindrical geometry that have been plotted as the remaining core thickness (x0  ) or radius (r0  ) (mm) versus immersion time t (h) for the noted acid concentrations. The average bone mineral content (BMC) of human cortical

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Figure 4 Remaining core thickness (x0  ) (mm) versus immersion time t (h) for human cortical bone for acid concentrations of () 0.5 N, () 1 N, and () 2 N HCl in planar geometry.

Figure 5 Remaining core radius (r0  ) (mm) versus immersion time t (h) for human cortical bone for acid concentrations of () 0.5 N, () 1 N, and () 2 N HCl in cylindrical geometry.

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bone determined from the residual weight of the bone discs upon complete demineralization was found to be 67  2%. The analytical models that allow prediction of demineralization depth as a function of immersion time and acid concentration are summarized briefly. For planar geometry, the depth of penetration of the demineralization process as a function of immersion time t in acid, (t) (cm) is given by (t) 

 2DC0t  103

(15)

where D is the effective mass diffusivity of HCl into the bone (cm2/s), C0 is the acid concentration (mol/L), t the immersion time in the HCl acid bath (s), and the moles of acid necessary to demineralize 1 cm3 of cortical bone completely (mol/cm3). The 103 factor is used to convert cubic centimeters to liters. Although the time dependence of the demineralization depth could also have been expressed analytically in cylindrical geometry, it would no longer have had the simple square root of immersion time dependence found in planar geometry. The time necessary to achieve a radial penetration depth of the demineralization process (t) with (t)  r0  rd (t) is given by 103 r 20 (r0  )2 r0   (r0  )2 2m   2  ln    1 2  0t()   2 4C0D r0 x r0 r 20









(16)

where r0 is the initial radius of a cylindrical bone and rd (t) the radial position of the demineralization front measured from the center line at time t. Note that the immersion time in hydrochloric acid required to obtain a certain demineralization depth was dependent on the initial radius of a cylindrical bone. The solution presented in Eq. (16) was therefore normalized for this dependence. E Fit of Experimental Data to Analytical Models In order to fit the experimental data to the analytical models, values of the effective mass diffusivities D at different acid concentrations and the value of as the amount of HCl required for complete demineralization of a specific volume of human cortical bone were determined. The value of was found by titration in HCl dilution series using bone discs of planar geometry. Titration experiments were carried out in triplicate and revealed to be 0.0144 mol/cm3. The effective mass diffusivities D of HCl at each normality in human cortical bone were determined using each , t pair and the model for planar geometry, as well as each , t pair and the model for cylindrical geometry, respectively. A student t test was employed to test for significance in the data comparing planar to cylindrical geometries. The mean values for D, the p values of the t test, and the maximum immersion time in HCl at each concentration for the two geometrical models are presented in Table 2. Kinetic data in planar and cylindrical geometry—when plotted as reduction of the dimensionless bone disc core thickness x*, with x*  x/x0 for planar geometry, or as the reduction of the dimensionless bone disc radius r*, with r*  rd/r0 for cylindrical geometry, versus the dimensionless immersion time t* with t*  Dt/x 20 for planar and t*  Dt/r 20 for cylindrical geometry on a log–log scale—fit well with the time dependencies predicted (Figs. 6 and 7). Linear regression yielded regression coefficients of 0.968  R2  0.999 of experimental data of planar geometry to the linear function predicted by Eq. (15). The mathematical models describing the demineralization kinetics of cortical bone in planar [Eq. (15)] and cylindrical geometry [Eq. (16)] were used to generate ready-to-use

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Table 2 Effective Mass Diffusivity of HCl in Demineralized Bone Matrix of Human Cortical Bone

CHCl (N) 0.5 1 2

Diffusivity of DBM of human cortical bone determined in planar geometry (106 cm2/s)

Diffusivity of DMB of human cortical determined in cylindrical geometry (106 cm2/s)

p values for D obtained from planar and cylindrical geometry

Maximum immersion time t for bone disc of planar geometry (h)

Maximum immersion time t for bone disc of cylindrical geometry (h)

2.31  0.39 2.35  0.39 1.33  0.20

2.69  0.52 2.50  0.29 0.92  0.15

0.180 0.038 1.7  107

17.3 8.5 7.5

46.5 25.0 34.0

Note: DMB, demineralized bone matrix.

reference tables that allow determination of the immersion time in HCl necessary to obtain a certain penetration depth of the demineralization process (see Tables 3 and 4). F Summary of Demineralization Studies These demineralization studies have demonstrated that hydrochloric acid demineralization of human cortical bone can be described with the theory of the shrinking core used for dif-

Figure 6 Demineralization depth in planar geometry plotted on a log–log scale as dimensionless bone disc core thickness x*  x/x0 versus the dimensionless immersion time t*  Dt/x 20 for HCl concentrations of () 0.5 N, () 1 N, and () 2 N. Data fit the t1/2 dependence given in Eq. (15) using values for the effective mass diffusivities D calculated from data in planar geometry and Eq. (15).

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Figure 7 Demineralization depth in cylindrical geometry plotted on a log–log scale as dimensionless bone disc core radius r*  rd/r0 versus the dimensionless immersion time t*  Dt/r 20 for HCl concentrations of () 0.5 N, () 1 N, and () 2 N. Data fit the time dependence predicted by Eq. (16) using values for the effective mass diffusivities D calculated from data in cylindrical geometry and Eq. (16). fusion limited processes in fluid–solid systems [25]. The shrinkage of the inner unreacted core is a function of the immersion time in acid. Cortical bone demineralization results in a sharp advancing reaction front that separates the demineralized bone matrix from an inner cortical bone core. As shown by the data (Figs. 6 and 7), the position of the reaction front can be accurately modeled for both planar and cylindrical geometry. The kinetic data in planar coordinates show that the penetration depth of the demineralization process increases linearly with the square root of the immersion time in hydrochloric acid. In cylindrical geometry, the kinetics is more complex due to the continuous change of curvature of the remaining inner cortical bone core. However, the demineralization process is insensitive to acid concentration higher than 1 N (see Table 2), where the diffusivity of the demineralized bone matrix is reduced for increasing acid concentrations. This decrease in the diffusivity results in a reduction of the speed at which the demineralization proceeds at higher acid concentrations. This effect is seen in the raw data (Figs. 4 and 5), where the difference in slope between the curves for 1 and 2 N is not as large as that between the 0.5 and 1.0 N curves. The reduction in diffusivity of demineralized cortical bone may be related to an alteration of collagen and other bone matrix proteins, which progresses with increasing immersion time and acid concentration. The protein degradation, as measured by nitrogen content of the acid liquor, has been found to be nearly constant for acid concentrations up to 4%w/v (1.1 N) HCl, but increases threefold when the concentration is increased to 8%w/v (2.2 N)

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Table 3 Demineralization Kinetics of Human Cortical Bone in Planar Geometry

Penetration depth (mm)

Immersion time in 0.5 N HCl (h:min)

Immersion time in 1 N HCl (h:min)

Immersion time in 2 N HCl (h:min)

0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 2.0 2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8 2.9 3.0

0 00:10 00:41 01:33 02:46 04:19 06:13 08:28 11:05 14:01 17:19 20:57 24:56 29:16 33:56 38:58 44:19 50:02 56:06 62:31 69:16 76:22 83:48 91:36 99:44 108:13 117:03 126:14 135:45 145:37 155:50

0 00:05 00:21 00:46 01:22 02:08 03:04 04:10 05:27 6:53 08:31 10:18 12:16 14:23 16:41 19:09 21:48 24:36 27:34 30:43 34:02 37:32 41:11 45:01 49:01 53:11 57:32 62:02 66:43 71:34 76:36

0 00:05 00:18 00:40 01:12 01:52 02:42 03:40 04:49 06:05 07:31 09:05 10:49 12:42 14:44 16:55 19:14 21:43 24:22 27:08 30:04 33:09 36:23 39:46 43:18 46:59 50:49 54:48 58:56 63:14 67:40

[29,33]. This results in an interdependence between the effective mass diffusivity D and the acid concentration C0. The reduced values for D at higher acid concentrations corroborate this hypothesis (Table 2). In addition, the D values obtained in planar and cylindrical geometry are significantly different for 1 and 2 N HCl (Table 2). The longer immersion time in cylindrical geometry for complete demineralization, as shown in Table 1, may have led to more extensive degradation of the bone matrix, thus reducing its diffusivity. Although transmission electron microscopy (TEM) studies have shown periodic striation of collagen fibers in rat bone matrix demineralized in 0.5 N HCl and in 2 N HCl suggesting that the architecture of the collagen network is not destroyed even at higher concentrations, it appears prudent to perform hydrochloric acid demineralization of cortical bone at low concentrations (1.0 N) to minimize protein degradation. The kinetics of HCl demineralization of human cortical bone in two geometries fre-

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Table 4 Demineralization Kinetics of Human Cortical Bone in Cylindrical Geometry Demine ralization depth (mm)

5

6

7

8

9

10

12

14

16

18

20

0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 2.0 2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8 2.9 3.0

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00:09 00:35 01:19 02:21 03:39 05:14 07:07 09:16 11:40 14:22 17:19 20:32 24:00 27:44 31:43 35:57 40:26 45:09 50:07 55:18 60:43 66:22 72:15 78:20 84:39 91:10 97:54 104:50 111:58 119:18

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00:09 00:35 01:20 02:22 03:40 05:17 07:10 09:20 11:47 14:31 17:31 20:47 24:20 28:09 32:14 36:34 41:11 46:02 51:10 56:32 62:10 68:02 74:10 80:32 87:08 93:59 101:04 108:59 115:56 123:43

00:09 00:35 01:20 02:22 03:41 05:17 07:11 09:22 11:49 14:34 17:34 20:52 24:26 28:17 32:23 36:46 41:25 46:19 51:29 56:55 62:37 68:33 74:45 81:12 87:53 94:50 102:02 109:28 117:08 125:02

00:09 00:35 01:20 02:22 03:41 05:17 07:11 09:22 11:50 14:35 17:37 20:56 24:31 28:22 32:31 36:55 41:35 46:32 51:44 57:13 62:57 68:56 75:12 81:43 88:29 95:29 102:46 110:17 118:03 126:04

00:29 00:35 01:20 02:22 03:41 05:18 07:12 09:23 11:52 14:37 17:40 20:59 24:35 28:27 32:36 37:02 41:44 46:42 51:56 57:27 63:13 69:16 75:34 82:07 88:56 96:01 103:21 110:56 118:47 126:52

Initial radius of a cylindrical bone graft (mm)

Note: Penetration depth versus immersion time (h:min) in 0.5 N HCl.

quently encountered in clinical bone allografts have been investigated. The planar model can be considered the best approximation for bone grafts with irregular planar shapes such as cortical struts. The cylindrical model is applicable for grafts with curved surfaces such as long diaphyseal bone grafts. The results of the model in both planar and cylindrical geometries have allowed prediction of the demineralization of cortical bone allografts of most sizes and shapes. They have provided the theoretical basis for modifying clinically used cortical bone allografts to grafts that have a highly osteoinductive surface layer. This new method of controlled demineralization should provide a means to produce cortical bone allografts that are rapidly incorporated and yet retain the biomechanical strength required for stable reconstruction. IV MECHANICAL PROPERTIES OF PARTIALLY DEMINERALIZED CORTICAL BONE GRAFTS The osteoinductive potential of demineralized cortical bone has been demonstrated in numerous experimental and clinical studies [9,12–14,34]. Demineralization of cortical bone

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results in exposure of osteoinductive bone matrix factors that enhance the incorporation of the graft [35–40]. However, the use of fully demineralized cortical bone allografts is not feasible in clinical practice where long bones must be reconstructed, because these grafts are not sufficiently strong to withstand normal skeletal forces. Stable fixation is practically impossible. Previous experiments on partially demineralized cortical bone plates obtained from dog femora demonstrated a moderate reduction of bending and compression strength with advancing demineralization [8]. With 81% demineralization of the test specimens, only a 42% loss of bending and a 24% loss of compression strength was observed. However, only limited conclusions can be drawn from this study because (1) percentage demineralization but not demineralization depth was related to the change of biomechanical properties, (2) cortical bone plates rather than whole bones were tested, and (3) an animal model instead of human bone was used. Therefore it has been anticipated to demonstrate the change of biomechanical properties of human diaphyseal bone allografts as a function of the demineralization depth. This study was designed to better define the feasible range of partial demineralization to produce diaphyseal bone allografts with enhanced osteoinductive and minimally altered biomechanical properties. As described, the demineralization process of cortical bone can be characterized as a diffusion rate limited process with a sharp advancing reaction front. In order to apply this process in clinical practice, it is also necessary to predict alterations in the biomechanical properties of cortical bone grafts following demineralization. Therefore, the changes in whole bone flexural rigidity (EI) in a model for human long bones (the fibula) in relation to the depth of the demineralization process was investigated. A Test Bones and Demineralization Human fibulae from young male donors were obtained from a bone bank that had discarded the bones because of bacterial contamination. The diaphyses of the test bones were stripped of soft tissue and cut to 15-cm lengths. Demineralization was carried out in a well-stirred 0.5 N hydrochloric acid (HCl) bath with an excess volume at room temperature. Test bones were controllably demineralized for varying immersion times using mathematical models allowing prediction of the demineralization depth as a function of the immersion time in HCl (Fig. 8). Bones were stored at 80°C until testing. Tested bones included four pairs (left and right fibula from the same donor) and 15 single fibulae. A total of 46 tests was performed. The four fibula pairs were tested twice in order to establish control tests, assess data reproducibility, and determine the range of normal left and right variability in flexural rigidity of human long bones utilizing the fibula as a representative model. The 15 single fibulae were divided, into five groups of three fibulae each including a control group with no demineralization (group 1), and four groups of test bones with increasing demineralization depth (group 2  0.2 mm; group 3  0.5 mm, group 4  0.75 mm; group 5  1.5 mm) (Fig. 8). Just prior to testing, bones were thawed in saline at room temperature. B Mechanical Testing The fibulae were tested in nondestructive four-point bending, following the polar flexural rigidity profile (PFRP) method described in detail by Foux et al. [41,42]. For clarity of presentation, the method is described briefly here. In order to determine the polar distribution of the flexural rigidity (EI) of a whole bone as a measure of its load-bearing capability, test bones were potted with low melting

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Figure 8 High resolution radiographs of 2-mm-thick cross-sections of partially demineralized fibulae excised from the center portion show advancement of the interface with increasing demineralization depth. These images were used to measure the inner (ri) and outer radii (r0) at the direction corresponding to where the stiffness index was determined. Demineralization depth: (a) 0.2 mm, (b) 0.5 mm, (c) 0.75 mm, and (d) 1.5 mm. Note the variations in size of the intramedulary canal and cortical thickness. The cortical thickness varied between 2 to 3 mm.

point bismuth alloy in two coaxial cups. The cups are an integral part of a four-point bending device and were mounted on an Instron testing machine. Cups and the potted test bone resembled a beam. This beam was rotated around its axis and locked at 15° angular increments. At each angular position, a low moment bending test (1.0 Nm) was conducted on the bone at a fixed rate of center deflection (0.5 mm/min) measured with a linear variable differential transformer (LVDT). A total of 25 bending tests was performed on each bone, with the last test repeating the first. The load and center deflection were recorded and used to plot load versus deflection curves. The linear portion of these curves was used to determine the slope by regression analysis. The flexural rigidity EI is the product of the modulus of elasticity and the moment of inertia at any cross-section of a bone. These properties have been assumed to be uniform along the longitudinal axis of the test bone. This assumption allowed for the calculations of representative EI values according to beam theory by using the slope of the load deflection curve and the test bone length. The 25 radial values for EI were thus obtained for each test, starting at the facies medialis of the fibula, encircling the entire perimeter of the bone in 15° angular increments and ending again at the medial direction. Consistency of the first and last test were used to test for data validity. The mean value of EI in the first and 25th test was taken as the EI value in the medial direction. Plotting the 24 EI values in polar coordinates formed an ellipse referred to as the PFRP. A typical PFRP plot is shown in Fig. 9. No absolute EI values were plotted because

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the objective was simply to demonstrate the change of the polar distribution of flexural rigidity by partial demineralization. The ellipses were curve fitted by least square regression that allowed for the definition of three parameters: major semiaxis, minor semiaxis, and , the angle of inclination. The flexural rigidity characteristics of each bone were thus defined by these three parameters. For comparison, each plot contains two PFRPs. The PFRPs of the four pairs of the left and right fibula were plotted together. For control bones of group 1 the PFRPs of the first and second test and for test bones of group 2 to 5 the PFRPs of each bone before and after demineralization were plotted together, respectively. Thus, if test results were identical, as one would expect for the control bones of group 1 tested twice, their PFRPs should coincide. In the case of group 2 to 5 test bones, differences will indicate the reduction in flexural rigidity due to demineralization. C Mechanical Parameters The ellipse parameters of the two PFRPs compared in one plot are used to define two mechanical parameters that are sensitive to changes in flexural rigidity. These included the stiffness index (SI) and the area ratio (AR). Stiffness index is defined as the ratio of the flexural rigidity EI after demineralization to before demineralization, both in the same direction that yields the minimum value for this ratio

 

EId SI   EIc

(17)

min

The area ratio is defined as the relative area of the EI ellipses that equals the square of the mean EIs: abd AR   abc

(18)

The numerical values of the mechanical parameters represent the biomechanical status of the partially demineralized bone to its status prior to demineralization. The SI and AR parameters therefore range between 1 and 0 or can also be considered as a percent difference between the two ellipses plotted in one PFRP. In the case of fibula pairs these ratios were calculated based on the ellipse parameters of left and right fibula, where the EId, abd were substituted with EIleft, ableft, and EIc, abc with EIright, abright, respectively. D Measuring of Test Bone Dimensions for Modeling In order to fit the experimental data for SI to the analytical model described, allowing for the prediction of the reduction in flexural rigidity as a function of the demineralization depth, it was necessary to obtain the exact dimensions of the cross-section of the test bones. Therefore, 2-mm-thick cross-sections were excised from the center portion of each test bone after testing. High resolution radiographs of these cross-sections were imaged and digitized using imaging software. The inner (ri) and outer (ro) radius were measured in 15° angular increments in the corresponding directions where the bending test was performed. E Control Fibula Pairs Four fibula pairs including the left and right bone from the same donor were tested twice each. The respective data are summarized as mean values of the two tests per pair in Table

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Table 5 Mechanical Parameters for Control Pairs Pair no.

SI

AR

1 2 3 4 Average

0.65 0.77 0.91 0.86 0.80  0.11

0.53 0.81 1.10 0.82 0.81  0.20

Note: Mean values of the first and second test.

5. The PFRP plots of these tests and the calculated mechanical parameters revealed considerable differences in flexural rigidity between normal left and right human fibula. The normal range of difference in rigidity between left and right fibula bone as expressed by SI and AR appeared to be as high as 35% for SI and 47% for AR, respectively (Table 3, pair no. 1; SI  0.65, AR  0.53). The average difference in rigidity between left and right test bone observed in the four fibula pairs was 20  11% for SI and 19  20% for AR, respectively (Table 5). In three out of the four pairs tested the right fibula was less rigid than the left. This finding could not be correlated with a preference for the left or right foot because no information about the handedness of the donors was available. All of the mentioned tests were repeated in order to assess test reliability, including reproducibility of orienting and potting the bones in the test device. Comparison of the experimental data for SI, AR, demin, and control from the first and second test of each pair revealed similar results. The maximum deviation of  between the first and second test was 11°. The same consistency in test reproducibility as assessed by comparing  before (control) and after demineralization (demin) was found in the tests of single fibulae of groups 1 to 5. F Single Fibulae The test results of single fibulae of groups 1 to 5 are summarized in Table 4. As expected, the ellipses of the PFRPs of group 1, the control fibulae, coincided almost exactly because these bones had not been treated with demineralization. As listed in Table 4, the SI and AR were approximately equal to 1. Figure 9 shows a typical PFRP plot for a control test bone of group 1 (bone no. 3; SI  0.97, AR  0.99). Test bones of group 2, which had been demineralized to a depth of 0.2 mm, revealed a reduction of SI  0.75 to 0.83 and AR  0.62 to 0.71. This reduction is within the range of normal variability in flexural rigidity between left and right human fibula. Figure 10 shows a typical PFRP plot of a fibula test bone demineralized to 0.2 mm deep. Further demineralization of the bones led to a drastic reduction of SI and AR with increasing demineralization depth as demonstrated by the mechanical parameters in test bones of group 3 to 5 (Table 6). G Analytical Modeling and Data Fit In order to derive a mathematical model for accurate prediction of the reduction of flexural rigidity in diaphyseal bone grafts following partial demineralization, their geometry was assumed to be a uniform hollow cylinder. A theoretical expression for SI has been derived and is presented in Eq. (19). The value of SI* is essentially the ratio of the moments of in-

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Figure 9 Polar flexural rigidity plot (PFRP) of a control test of a human fibula (test bone no. 3, Table 4) without demineralization. () First test (experimental EI; solid line, regression fit). () Second test (experimental EI; dashed line, regression fit). SI  0.9719; AR  0.9969, first test (•)  238°; second test ()  244°.

ertia of two hollow cylinders with the same inner but a different outer diameter due to demineralization. rd4ƒ  ri4 SI*    r4o  ri4

(19)

where ro is the outer radius of the graft, ri the inner radius of the graft and rdƒ the radial position of the demineralization front. Taking the values of ri and ro obtained from high resolution x-ray image analysis at that angular position where SI was calculated allowed for the experimental SI values for each bone to be fitted to the analytical model presented in Eq. (19). In order to simplify plotting the best fit lines of the experimental SI to Eq. (19), the relation was normalized for ro as

r  r . SI*   r 1   r rdƒ

ri

4

o

o

ro

i

4

4

(20)

o

The reduction of cortical thickness following demineralization was expressed in the demineralization index (DI), which takes the inner and outer radius of a diaphyseal bone allograft into account. The relation for DI was also normalized for ro in order to clearly demonstrate the relation between demineralization depth and reduction of flexural rigidity of diaphyseal bone grafts when plotting SI* versus DI.

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Figure 10 Polar flexural rigidity plot of a human fibula partially demineralized (test bone no. 4, Table 4). () Before demineralization (experimental EI; solid line, regression fit). () After demineralization (experimental EI; dashed line, regression fit). SI  0.8321; AR  0.7110; control (•)  219°; demin ().

Table 6 Mechanical Parameters for Test Bones in Groups 1 to 5 Group 1

2

3

4

5

Bone no.

Demineralization depth (mm)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15

Control bones, no demineralization 0.20

0.50

0.75

1.50

SI

AR

0.9619 1.0021 0.9719 0.8321 0.7729 0.7525 0.5865 0.4854 0.5095 0.3928 0.4027 0.5882 0.0657 0.0671 0.0592

0.9456 1.0956 0.9969 0.7110 0.6809 0.6298 0.3769 0.3001 0.3007 0.1827 0.1848 0.3810 0.0069 0.0072 0.0042

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Figure 11 Fit of experimental SI to analytical model of a hollow cylinder plotted as SI* versus DI as a dimensionless variable for the demineralization depth. Data correlated well with the r4 dependence predicted by Eq. (20) and (21). Included are curves of the projected SI* reduction based on the mean, minimum, and maximum ri/ro ratios observed for the test bones.

r  r . DI   r 1   r rdƒ

ri

o

o

(21)

i

o

The experimental and theoretically expected SI [Eq. (19)] are plotted in Fig. 11 versus demineralization depth expressed as the demineralization index DI [Eq. (21)]. The solid line in Fig. 4 represents the reduction of SI with increasing demineralization depth for the mean ri/ro ratio observed in the test bones. The dashed lines represent the projected SI reduction for the minimum and maximum ri/ro ratio of the test bones. The experimental SI data are located in the range defined by the lines of the minimum and maximum ri/ro ratio and followed the r4 dependence predicted by Eq. (19). V MECHANICAL PROPERTIES OF PERFORATED AND DEMINERALIZED BONE GRAFTS The purpose of this study was to extend previous biomechanical measurements of demineralized bone to grafts that have been first perforated with the use of an Er:YAG laser and then additionally surface decalcified. Knowledge of such subsequent alterations of mechanical properties is required before these types of grafts can be considered for clinical use. Ideally, improvement of osteoinductive properties can be achieved while only mini-

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mally altering the strength of the original bone. In order to verify this hypothesis, the impact of laser perforation and partial demineralization on flexural rigidity and compression strength of diaphyseal bones has been investigated. A Test Bones and Preparation Eighteen tibiae were harvested from 1-year-old sheep. The diaphysis of the test bones were stripped of soft tissue and cut to 12-cm lengths. An erbium:yttrium-aluminum-garnet laser (Schwartz Electro-Optics, Concord, MA) operating at a wavelength of 2940 nm was used to drill holes of 300 m in diameter through the entire thickness of the diaphyseal cortex producing a hole that was typically 2 mm deep. The laser was optically connected to an OPMI 1 F/C operating microscope (Zeiss, Oberköchen, Germany) by an articulated arm. A 100–200 s pulse train consisting of pulses g1s long was used for drilling. The energy delivered per pulse was measured using a pyroelectric joulemeter (ED-200, Gentec, Ste. Foy, Quebec) and oscilloscope. A collinear helium–neon laser with a visible beam was used to aim the Er:YAG laser. The beam diameter at the tissue surface was 330 m as defined by the e2 intensity points. The lasers were aimed using a micromanipulator (LaserMechanisms, Southfield, MI). Typically, drilling was performed employing a 53-mJ pulse, with 25–30 pulses required for the drilling of each hole. The fluency per pulse was typically 60 J/cm2. Only a 5-cm-long center section of bone diaphyses was treated by laser hole drilling. As described, this is equal to the exposed length of the bone in the bending test device. That is, no holes were drilled in the remaining outer portions of each end. Holes were drilled in rows of 25 holes each around the circumference of the diaphysis such that the distance between holes within each row and to the next row was 2 mm allowing for the generation of a uniform hole grid. Typically, this hole grid consisted of 300 holes. For demineralization, the bone ends were capped with an acid-resistant paraffin layer thereby allowing demineralization to proceed only from the periosteal surface to the endosteal surface. Demineralization was accomplished by immersing the bones in a wellstirred volume excess 0.5 N hydrochloric acid bath at room temperature. The test bones were demineralized in a controlled manner using mathematical models to determine the demineralization depth as a function of the immersion time in hydrochloric acid. These models describing the demineralization kinetics accounted for variances in the bone size as given by the outer radius. The extent of demineralization was chosen to produce a uniform 20% reduction of the rigidity of the original bone. After demineralization, test bones would therefore be left with 80% of the rigidity of the original bone, provided that no other treatment that could further reduce their rigidity, i.e., laser drilling, was instituted. Bones were stored at 80°C until testing. Tested bones included 18 tibiae from nine donor animals. These were divided into three groups of six tibiae each: a control group having no demineralization or laser drilling (group 1), a second group having laser drilling only (group 2), and a third group having laser drilling and demineralization (group 3). In this study, the immersion time was normalized by the outer radius of the test bone. All bones were thawed in normal saline at room temperature just prior to testing. B Mechanical Testing The sheep tibiae were tested in two different settings. First, a nondestructive four-point bending test following the polar flexural rigidity profile method, described in detail by

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Foux et al. [41,42], was performed. Second, bone specimens were subsequently tested for their compression strengths by loading them to failure in the axial direction. For clarity of presentation, both methods are described here. For nondestructive four-point bending, bone ends were potted with low melting point bismuth alloy in two coaxial cups. The cups and the potted test bone formed a beam that was mounted on a four-point bending apparatus in an UTS 10 (Ulm, Germany) testing machine. This beam was capable of being rotated about its longitudinal axis and locked at 15° angular increments starting at the anterior tuberosity of the tibial bone. At each subsequent angular position, a low moment bending test (1 N-m) was conducted on the bone at a fixed rate of center deflection (0.5 mm/min). The center deflection was measured with a linear variable differential transformer. Twenty-five bending tests were performed on each bone, with the last test repeating the first. The load and center deflection were recorded and used to plot a load versus deflection curve for each bending test. The slope was calculated from the linear portion of these curves by regression analysis. The flexural rigidity was found as the product of the modulus of elasticity and the moment of inertia at any cross-section of a bone. Assuming the flexural rigidity to be uniform along the longitudinal axis of the test bone, beam theory was used to calculate representative values from the slopes of the load–deflection curves and the exposed length of test bones (52 mm). For each bone, 25 elliptically distributed values for the flexural rigidity were obtained. The repeatability of the first and last test was used as a measure of the reliability of the data. The mean value from the first and 25th test was taken as the rigidity value in the anterior direction. The 25 flexural rigidity values from each bone were plotted in polar coordinates. The absolute values have not been included because only relative changes in flexural rigidity were assessed. An ellipse was curve fitted to the measured flexural rigidity values of each bone by least square regression to generate a polar flexural rigidity profile. Three parameters were obtained from each of these profiles: major semiaxis (a), minor semiaxis (b), and the angle of inclination (). The flexural rigidity characteristics were defined by these three parameters. Each plot contains two polar flexural rigidity profiles for comparison of the before and after states of each individual bone subjected to a specific treatment (groups 1–3). Each bone served as its own control. Thus, if test results were identical, as one would expect in group 1, their polar flexural rigidity profiles should coincide. In groups 2 and 3, differences will indicate the reduction in flexural rigidity caused by laser drilling and demineralization. Following the nondestructive four-point bending test, bone specimens were cut to 3cm-long cylinders. These were taken from the center section of the diaphyseal portion such that the entire length of the cylinder was either laser drilled (group 2) or laser drilled and demineralized (group 3). Bone cylinder specimens of all three groups were then tested for their compression strength by applying axial load to failure in a UTS 10 testing machine at a cross head speed of 0.5 mm/min. Data on the failure loads of the bone cylinder specimens were recorded and used for comparative statistical analysis of failure strength between bone cylinder specimens of groups 1–3 by employing a one-way analysis of variance (ANOVA). The stress of the load–deflection curve could be calculated by dividing the applied load by the cross-sectional area that had been obtained from measurements of bone dimensions. It was assumed that the diaphyseal bone specimens were of an ideal hollow cylinder geometry. Similarly, displacement of the bone specimen was converted into strain by dividing by the length of the original specimen.

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C Mechanical Parameters For interpretation of results of the nondestructive four-point bending test, the ellipse parameters of the two flexural rigidity profiles compared in each plot were used to define two mechanical parameters that are sensitive to changes in test bone flexural rigidity. These parameters include the stiffness index and the area ratio. The stiffness index (SI) is defined as the ratio of the flexural rigidity (EI) values after and before treatment, both in the same direction. This yields the minimum value of this ratio:



EIafter treatment SI   EIbefore treatment



min

The area ratio (AR) is defined as the relative area of the ellipse equal to the square of the ratio of the mean flexural rigidity values AR  abafter treatment/abbefore treatment. The stiffness index is the most meaningful parameter because it is directly related to the moment of inertia. The numerical values of the mechanical parameters represent the biomechanical status of a laser drilled (group 2) or laser drilled and partially demineralized diaphyseal bone graft (group 3) relative to its status before treatment. Values of the stiffness index and the area ratio range between 1 and 0; a value of 1 indicates no change in the flexural rigidity, and a value of 1 indicates a reduction in the flexural rigidity of a bone as a result of laser drilling and demineralization. For the compression experiment, a percentage strength was calculated using the data of the control group as the baseline values. D Measuring of Test Bone Dimensions Before test bones were subjected to treatment, the outer radius had been measured and recorded. This allowed an immersion time corrected for variances in the outer radius of each bone to be determined. These measurements were made at the center section of the tibia at the 0 and 90° angular positions. The mean of these values was then used as the outer radius for each test bone based on the assumption of cylindrical geometry. The cortical thickness was obtained from radiographs of the center section of the tibia at the 0 and 90° angular positions permitting the calculation of the inner radius as the mean of these measurements. These parameters were used to calculate the cross-sectional surface area at the center section. E Results of Bending Test Control test bones of group 1 were tested twice to provide polar flexural rigidity profile plots of diaphyseal bones that had not undergone treatment. As expected, polar flexural rigidity profiles of all bones of the control group coincided almost perfectly. As listed in Table 7, the values for stiffness index and area ratio were very close to 1. Also, similar values of  were found in the first and second test indicating good reproducibility of results of the test method. The maximum deviation of  has been noted to be 6°. Laser perforation of the bones in group 2 caused an average reduction in flexural rigidity that resulted in the following: SI  0.965  0.034 (mean  standard deviation), and AR  0.965  0.067 (Table 5). Figure 12 shows a typical polar flexural rigidity profile plot for a laser-perforated bone. Group 3, containing bone samples that had been laser perforated and then partially

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demineralized, showed a drastic reduction in flexural rigidity as indicated by a mean stiffness index of 0.595  0.066 and a mean area ratio of 0.375  0.0790 (Table 7). A typical polar flexural rigidity profile plot is given in Fig. 13 for a laser-perforated and partially demineralized bone. Upon statistical analysis by ANOVA testing, no significant difference in stiffness index and area ratio between control (group 1) and laser-perforated (group 2) test bones was found. However, the same test revealed a significant reduction of rigidity for laser-perforated and partially demineralized test bones (p  2.6  107) when compared with the control group. F Results of Compression Testing Cylindrical center sections of all test bones used in the nondestructive bending test were loaded to failure. The respective compression strength data are listed in Table 8. In laser-perforated bone specimens (group 2), the percentage strength was reduced by 2.7%. Additional partial demineralization (group 3) caused a percentage strength reduction of 12.7%. Again, ANOVA testing revealed the percentage strength of laser-perforated and partially demineralized test bones to be significantly lower as compared to bone specimens of groups 1 and 2. As shown in Fig. 14, the average stress–strain curves for the control (group 1) and laser-perforated (group 2) bone specimens are characterized by comparable slopes, maximum stress, and strain at the yield point. The average stress–strain curve of group 3 test bones can be divided into two parts. The initial period is characterized by an elastic behavior after which deformation occurs in the plastic mode in the second part of the curve with a lower ultimate stress at the yield point.

Table 7 Mechanical Parameters of Bending Test in Groups 1 to 3 Group 1

2

3

Bone no. 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18

               

Demineralization depth (m)

None

None

337.5 336.8 392.3 355.3 385.1 318.4

Mean stiffness index

SI 0.98 0.96 0.99 1.00 0.97 0.98 0.99 0.99 1.00 0.96 0.92 0.93 0.63 0.71 0.52 0.57 0.56 0.58

                       

0.980  0.0141

0.965  0.034

0.595  0.006

Mean area ratio

AR 0.99 0.95 0.97 1.10 1.00 0.99 1.01 1.02 1.04 0.94 0.89 0.89 0.42 0.51 0.29 0.34 0.33 0.36

                       

1.000  0.052

0.965  0.067

0.375  0.079

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Figure 12 Polar flexural rigidity profile of a laser-perforated test bone without partial demineralization (bone 10, Table 4). () Test before laser perforation (experimental flexural rigidity; solid line, regression fit). (*) Test after laser perforation (experimental flexural rigidity; dashed line, regression fit). SI  0.96; AR  0.94; first test  254.2°; second test  254.79°.

VI DISCUSSION Alteration of surface geometry of demineralized cortical bone by laser perforation has been found to promote bony ingrowth leading to improved osteogenesis and enhanced incorporation of these types of bone grafts [9,37,43,44]. However, before their clinical application in the reconstruction of long bones, evaluation of their biomechanical properties is necessary. The elegant PFRP test method and analysis produces quantitative measures for defining the mechanical status of a diaphyseal bone. It is based on a four-point bending test that does not destroy the integrity of the test bone. By determining the polar distribution of the flexural rigidity (EI) in 15° angular increments, it has become possible to define mechanical parameters that are extremely sensitive to changes in the rigidity and the asymmetry of a partially demineralized bone. The graphic presentation produced by superimposing the two PFRPs of a bone tested before and after partial demineralization provides an easy-tounderstand means for immediate comparison and determination of the biomechanical status of diaphyseal bones in relation to the demineralization depth. The use of the polar flexural rigidity profile test method and analysis in this study was advantageous for several reasons: It has produced quantitative measures of the mechanical

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Figure 13 Polar flexural rigidity profile of a laser-perforated and partially demineralized test bone (bone 18, Table 4). () Test before laser perforation and partial demineralization (experimental flexural rigidity; solid line, regression fit). (*) Test after laser perforation and partial demineralization (experimental flexural rigidity; dashed line, regression fit). SI  0.58; AR  0.36; first test  244.81°; second test  243.94°.

status of whole bones. The method has been previously used for quantitative measures of mechanical properties of whole experimental bones and bone grafts and is therefore suitable for comparison [28]. Although reduction of flexural rigidity and ultimate failure strength have been found empirically to be highly correlated, no direct relationship exists [41,45]. Testing for compression strength was therefore useful in determining absolute values and differences in ultimate failure strength between the three experimental groups. Table 8 Results of Compression Test in Groups 1 to 3

Group

Number of specimens

Mean failure load (kN)

Loss of compression strength (%)

Maximum displacement (mm)

1 2 3

6 6 6

10.82 10.59 09.43

02.7 12.7

0.450 0.473 0.799

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Figure 14 Average stress–strain curve for control bones (bold solid line: group 1), laser-perforated bones (thin solid line: group 2), and laser-perforated and partially demineralized bones (dashed line: group 3). Preparation of test bones was performed with the use of highly reproducible techniques for both laser perforation and demineralization. As previously noted, demineralization of cortical bone is extremely predictable following the kinetics of a diffusion rate–limited process. The Er:YAG laser has allowed for the removal of bony tissue by explosive vaporization thus eliminating the need for mechanical contact with the tissue [24,46]. Holes of reproducible diameter could be drilled without destroying the integrity of the tissue surrounding the holes. Precise drilling was made possible by the short optical penetration depth of the 2.94 m mid-infrared pulsed laser radiation [25]. Thermal damage, with subsequent denaturation or carbonization of tissue, was minimized because heat deposition is confined to the volume heated by the laser [46]. Bone grafts with a homogeneous laser hole grid and a constant distance between holes can therefore be constructed. Therefore, inconsistency in test results because of variations in either laser drilling or demineralization was not expected. This is supported by the small standard deviations of mechanical parameters listed in Tables 4 and 6. In the first set of experiments on demineralization alone, repeated testing of fibula pairs revealed a high degree of reproducibility and consistency between the different tests. The minor differences in SI and AR observed between the first and second tests are most likely related to variations in the orientation and potting of the bones in the test device as suggested by the differences in . Testing bone pairs also facilitated the determination of the normal range of variability in flexural rigidity between left and right human long bones using the fibula as a representative model. The stiffness index SI [Eq. (17)] appears to be the most meaningful parameter in assessing the variability in flexural rigidity because it is

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directly related to the moment of inertia of a cross-section of a diaphyseal bone. The flexural rigidity of diaphyseal bones as assessed by the stiffness index SI may vary considerably between left and right bones as impressively demonstrated by these data. Variations from 15–35% between the left and right bones appear to be the normal range. A similar interpersonal variability has been reported on the large series of human bones [47]. Standard deviations of 15% for the elastic modulus in the axial direction and 30% for Poisson’s coefficient have been found. These data on interpersonal variability of flexural rigidity may be helpful in determining the depth of demineralization in diaphyseal bone grafts necessary for maximum osteoinductive properties and yet leaving a comfortable safety margin allowing the graft to withstand normal skeletal forces. Test results in single fibulae of groups 2 to 5 revealed that SI and AR were dramatically reduced with increasing demineralization depth. The flexural rigidity of diaphyseal bones as assessed by SI and AR has proven to be highly dependent on changes in geometry, in particular, the outer radius due to partial demineralization. This relation is expressed in the analytical model presented in Eq. (19) that reveals a fourth power dependence of SI* on the outer radius of a diaphyseal bone. Although the tested fibulas were not entirely cylindrical (Fig. 1), the experimental data fitted the ideal hollow cylinder model surprisingly well (Fig. 4). The deviation of the experimental SI data from the analytical model [Eq. (20) and (21)] can be primarily explained by the difference between the noncylindrical geometry of the test bones and the ideal hollow cylinder model. In addition, the SI per definition was based solely on EI values taken at one polar angle on the ellipses, neglecting test results from other directions. In the second series of experiments on perforated and partially demineralized test bones, results of the nondestructive four-point bending test also showed coinciding polar flexural rigidity plots for the control bone samples of group 1. This was expected because these bones were neither laser perforated, nor demineralized, and the polar flexural rigidity profile was generated by repeating the original test on the same sample. Minor variations were noted in , indicating a high degree of reproducibility between different tests, and were thought to be most likely related to variations in orienting the bones in the test device. Laser perforation resulted in insignificant changes of flexural rigidity. However, large reductions were noted when laser drilling was combined with partial demineralization. This finding may be attributable to the fact that the removal of relatively small amounts of bony tissue by laser hole drilling does not substantially change the anatomical dimensions of the original bone. This, however, does occur during demineralization, where the inner bone mineral–containing portion shrinks continuously with increasing demineralization depth [4]. Considering that the latter parameter is directly related to the stiffness index by a fourth power dependence [28], larger reductions of rigidity are expected following partial demineralization than laser perforation. Results of the compression testing corroborated findings of the nondestructive fourpoint bending test. While laser perforation alone resulted in an almost negligible reduction, much smaller values for the compression strength were found when additional demineralization was employed. However, the percentage loss of compression strength was only onequarter of the reduction of flexural rigidity. Analysis of the average stress–strain curves revealed pure linear behavior in control and laser-perforated bones. For the laser-perforated and partially demineralized bones, the initial part of the stress–strain curve shows an elastic behavior with essentially all of the increased displacement taking place in this part of the curve. This behavior is consistent with the elastic/perfectly plastic model reported for

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bone tissue [38]. This model implies that bone behaves as a “two-phase” material consisting of mineral and matrix, where the mineral generates most of its elastic stiffness and the collagen most of its plastic stiffness. Moreover, the experimental data reflect this behavior because most of the deformation occurs in the plastic mode with similar slopes in all three experimental groups. VII CONCLUSIONS This study addressed the issue of biomechanical feasibility of partial demineralization for clinically used diaphyseal bone allografts. It is unique because (1) human bones representative of the clinical situation were tested, (2) the graphic presentation of PFRP results have unquestionably demonstrated changes in the biomechanical properties of diaphyseal bone grafts in relation to the demineralization depth, and (3) an analytical model has been presented that allows prediction of the stiffness index of diaphyseal bones independent of their size as a function of the demineralization depth. This study has shown that only shallow demineralization appears to be feasible for applicability of partial demineralization for clinically used diaphyseal bone allografts. Taking the range of normal variability in flexural rigidity between left and right human fibula bones into account, partial demineralization resulting in a reduction of the stiffness index of 20–30% appears to be the maximum acceptable extent. This reduction, as expressed in the analytical model [Eq. (20) and (21)], is not only a function of the demineralization depth, but is also dependent on the initial inner and outer radius. Thus, the desirable demineralization depth may vary among diaphyseal bones depending on their cortical thickness and the size of the intramedullary canal. The findings of this study suggest that the simple hollow cylinder model can be used to predict the biomechanical properties of diaphyseal bone allografts following demineralization. Furthermore, complex testing more closely simulating the clinical situation after implantation of partially demineralized diaphyseal bone allografts may not be necessary as long as the demineralization extent does not exceed the inter- and intrapersonal variability of rigidity. The ultimately feasible extent of demineralization in laser-perforated diaphyseal bone grafts was evaluated in the second mechanical study in order to provide additional mechanical parameters for their safe clinical use. Shallow demineralization resulting in a reduction of the stiffness index of 20 to 30% was found to be the acceptable range based on the previous mechanical tests on partially demineralized cortical bone grafts [28]. A fixed limit of demineralization equivalent to a 20% reduction of the stiffness index was therefore chosen in the contingency study because additional weakening of the graft was expected due to the laser drilling. It appears noteworthy that the observed reduction in the stiffness index of the laserperforated and partially demineralized group was larger than the expected additive effect of each single treatment. Instead of the predicted 20 to 25%, an average reduction of rigidity of 40.5% was found. This suggests that more complex changes in the composite structure of bone are introduced by both the laser perforation and the demineralization process. Even though demineralization leads to an increasingly elastic behavior [26], a dramatic reduction of the stress at the yield point of laser-perforated and partially demineralized bone specimens appears evident. Taking all of these considerations together, controlled shallow partial demineralization and perforation of a diaphyseal bone allograft appears to be feasible utilizing the math-

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ematical models presented herein and should result in a graft with improved osteoinductive and sufficient biomechanical properties.

ACKNOWLEDGMENTS This work was supported by NIH Grant AR-21896, the Medical Free Electron Laser Program under Office of Naval Research Contract N00014-91-0084, and by DOE Grant # DEFG02-91 ER61228, and NIH/NIAMS grant AR 45062 (to Kai-Uwe Lewandrowski). No ownership or profit has been derived from this study. REFERENCES 1. Aro H. T., Aho A. J. 1993. Clinical use of bone allografts. Ann. Med. 25:403–412. 2. Berrey B. H. J., Lord C. F., Gebhardt M. C., Mankin H. J. 1990. Fractures of allografts: frequency, treatment, and end-results. J. Bone Joint Surg. Am. 72:825–833. 3. Mulliken J. B., Glowacki J., Kaban L. B., Folkman J., Murray J. E. 1981. Use of demineralized allogeneic bone implants for the correction of maxillocraniofacial deformities. Ann. Surg. 194:366–372. 4. Pak J. H., Paprosky W. G., Jablonsky W. S., Lawrence J. M. 1993. Femoral strut allografts in cementless revision total hip arthroplasty. Clin. Orthop. 295:1–6. 5. Wozney J. M. 1989. Bone morphogenetic proteins. Prog. Growth Factor Res. 1:267. 6. Tomford W. W., Starkweather R. J., Goldman M. H. 1981. A study of the clinical incidence of infection in the use of banked allograft bone. J. Bone Joint Surg. 63A:244–248. 7. Tomford W. W., Thongphasuk J., Mankin H. J., Ferraro M. J. 1990. Frozen musculoskeletal allografts. A study of the clinical incidence and causes of infection associated with their use. J. Bone Joint Surg. 72A:1137–1143. 8. Guo M. Z., Xia Z. S., Lin L. B. 1991. The mechanical and biological properties of demineralized cortical bone allografts in animals. J. Bone Joint Surg. 73B:791–794. 9. Hosny M., Sharawy M. 1985. Osteoinduction in young and old rats using demineralized bone powder allografts. J. Oral Maxillofac. Surg. 43:925–931. 10. Narang R., Wells H., Laskin D. M. 1982. Experimental osteogenesis with demineralized allogeneic bone matrix in extraskeletal sites. J. Oral Maxillofac. Surg. 40:133–141. 11. Vandersteenhoven J. J., Spector M. 1983. Histological investigation of bone induction by demineralized allogeneic bone matrix: a natural biomaterial for osseous reconstruction. J. Biomed. Mater. Res. 17:1003–1014. 12. Glowacki J., Mulliken J. B. 1985. Demineralized bone implants. Clin. Plast. Surg. 12:233–241. 13. Kaban L. B., Mulliken J. B., Glowacki J. 1982. Treatment of jaw defects with demineralized bone implants. J. Oral Maxillofac. Surg. 40:623–626. 14. Mulliken J. B., Glowacki J., Kaban L. B., Folkman J., Murray J. E. 1981. Use of demineralized allogeneic bone implants for the correction of maxillocraniofacial deformities. Ann. Surg. 194:366–372. 15. Sayler K. E., Gendler E., Menendez J. L., Simon T. R., Kelly K. M., Bardach J. 1992. Demineralized perforated bone implants in craniofacial surgery. J. Craniofac. Surg. 3:55–62. 16. Tyler H. D., Huckstep R. L., Stalley P. D. 1987. Intraluminal allograft restoration of the upper femur in failed total hip arthroplasty. Clin. Orthop. 224:26–32. 17. Salyer K. E., Bardach J., Squier C. A., Gendler E., Kelly K. M. 1995. Cranioplasty in the growing canine skull using demineralized perforated bone. Plast. Reconstr. Surg. 96:770–779. 18. Mankin H. J., Gebhardt M. C., Jennings L. C., Springfield D. S., Tomford W. W. 1996. Longterm results of allograft replacement in the management of bone tumors. Clin. Orthop. 324:86– 97.

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19. Jofe M. H., Gebhardt M. C., Tomford W. W., Mankin H. J. 1988. Reconstruction for defects of the proximal part of the femur using allograft arthroplasty. J. Bone Joint Surg. 70A:507–516. 20. Enneking W. F., Mindell E. R. 1991. Observations on massive retrieved human allografts. J. Bone Joint Surg. 73A:1123–1142. 21. Mankin H. J., Springfield D. S., Gebhardt M. C., Tomford W. W. 1992. Current status of allografting for bone tumors. Orthopedics 15:1147–1154. 22. Cummings J. P., Walsh Jr. J. T. 1993. Erbium laser ablation: effects of dynamic optical properties. Appl. Phys. Lett. 32:493–503. 23. Schomacker K. T. 1996. Effect of Er:YAG laser holes on osteoinduction in demineralized rat calvarial allografts. J. Orthop. Res. 14:108–113. 24. Walsh Jr. J. T., Deutsch T. F. 1989. Er:YAG laser ablation of tissue: measurement of ablation rates. Lasers Surg. Med. 9:327–337. 25. O’Donnell R. J., Deutsch T. F., Flotte R. J., Lorente C. A., Tomford W. W., Mankin H. J., Schomacker K. T. 1996. Effect of Er:YAG laser holes on osteoinduction in demineralized rat calvarial allografts. J. Orthop. Res. 14:108–113. 26. Arends J., Christoffersen J., Christoffersen M. R., Ogaard B., Dijkman A. G., Jongebloed W. L. 1992. Rate and mechanism of enamel demineralization in situ. Caries Res. 26:18–21. 27. Arends J., Dijkman T., Christoffersen J. 1987. Average mineral loss in dental enamel during demineralization. Caries Res. 21:249–254. 28. Lewandrowski K.-U., Venugopalan V., Tomford W. W., Schomacker K. T., Mankin H. J., Deutsch T. F. 1996. Kinetics of cortical bone demineralization. A new method for modifying cortical bone allografts. J. Biomater. Res. 31:365–372. 29. Makarewicz P. J., Harasta L., Webb S. L. 1980. Kinetics of acid diffusion and demineralization of bone. J. Photog. Sci. 22:148–159. 30. Birkedal-Hansen H. 1974. Kinetics of acid demineralization in histologic technique. J. Histochem. Cytochem. 22:434–441. 31. Levenspiel O. 1972. Chemical Reaction Engineering. Wiley & Sons: New York, pp. 364–365. 32. Paprosky W. G., Perona P. G., Lawrence J. M. 1994. Acetabular defect classification and surgical reconstruction in revision arthroplasty. A 6-year follow-up evaluation. J. Arthroplasty 9:33–44. 33. Makarewicz J., Harasta L., Webb S. L. 1981. Application of shrinking core kinetics to bone demineralization. AICHE. Symp. Ser. 202(77):141–149. 34. Upton J., Glowacki J. 1992. Hand reconstruction with allograft demineralized bone: twenty-six implants in twelve patients. J. Hand Surg. Am. 17:704–713. 35. Urist M. R., DeLange R. J., Finerman G. A. 1983. Bone cell differentiation and growth factors. Science 220:680–686. 36. Wozney J. M. 1989. Bone morphogenetic proteins. Prog. Growth Factor Res. 1:267–280. 37. Young M. F., Kerr J. M., Ibaraki K., Heegaard A. M., Robey P. G. 1992. Structure, expression, and regulation of the major noncollagenous matrix proteins of bone. Clin. Orthop. 281:275–294. 38. Boskey A. L. 1992. Mineral–matrix interactions in bone and cartilage [Review]. Clin. Orthop. 281:244–274. 39. Centrella M., Horowitz M. C., Wozney J. M., McCarthy T. L. 1994. Transforming growth factor-beta gene family members and bone [Review]. End. Rev. 15:27–39. 40. Aldinger G., Herr G., Kusswetter W., Reis H. J., Thielemann F. W., Holz U. 1991. Bone morphogenetic protein: a review. Int. Orthop. 15:169. 41. Foux A., Black R. C., Uhthoff H. K. 1990. Quantitative measures for fracture healing: an invitro biomechanical study. J. Biomech. Eng. 112:401–406. 42. Foux A., Uhthoff H. K., Black R. C. 1993. Healing of plated femoral osteotomies in dogs. A mechanical study using a new test method. Acta Orthop. Scand. 64:345–353. 43. Salyer K. E., Gendler E., Menendez J. L., Simon T. R., Kelly K. M., Bardach J. 1992. Demineralized perforated bone implants in craniofacial surgery. J. Craniofac. Surg. 3:55–62.

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23 Adenovirus Vector–Mediated Gene Transduction for the Treatment of Bone and Joint Destruction of Rheumatoid Arthritis Sakae Tanaka The University of Tokyo, Tokyo, Japan

I INTRODUCTION Rheumatoid arthritis (RA) is a chronic inflammatory disorder characterized by synovial hyperplasia [1]. Investigation into the pathogenesis of joint destruction of RA has revealed the transformed phenotype of rheumatoid synovial cells. Proliferation of the synovial cells leads to pannus formation that invades the bare area between cartilage and bone, and causes progressive bone erosion and periarticular osteopenia in the affected joints. Radiographic studies demonstrate that bone erosion of RA begins early in the disease and progresses throughout the course of the disease. Bone erosion results in severe deformity of the affected joints and impairs the normal activity of the patients. Therefore, inhibiting the bone destruction in RA is one of the most important issues in treatment. Because the exact etiology of RA remains unknown, most treatments are just treating symptoms of the disease. Nonsteroidal anti-inflammatory drugs (NSAIDs) have been used to reduce the painful symptoms of the disease, but they have little effect on stopping the progression of the joint destruction. Some disease-modifying antirheumatic drugs (DMARDs) reportedly reduce the inflammation and suppress joint destruction in RA. However, in most cases bone-protective effects of these agents are limited, and prolonged usage of them sometimes causes severe side effects of stomach ulceration and hepatotoxicity. There is accumulating evidence that osteoclasts, primary cells responsible for bone resorption, are involved in the bone destruction in RA, and recent progress in molecular biology and biochemistry has 467

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revealed the molecular mechanism of osteoclast differentiation and bone resorption [2]. In this chapter, I would like to focus on the role of osteoclasts in bone and joint destruction in RA and propose that they can be potential targets of RA therapy. We also found that adenovirus vectors can efficiently transduce osteoclasts, and by modulating intracellular signaling pathways in osteoclasts using adenovirus vectors, we could successfully regulate their function not only in vitro but also in vivo. II STRUCTURES AND FUNCTIONS OF OSTEOCLASTS Osteoclasts are multinucleated giant cells (50–100 m average) derived from hematopoietic precursor cells of monocyte–macrophage lineage, and highly differentiated for bone resorption (Fig. 1). Osteoclasts are highly motile cells, and once attached on the bone surface, the cells become highly polarized and form a tight ringlike zone of adhesion, the sealing zone [3]. The cytoplasm of the sealing zone is devoid of cellular organelles, except for numerous actin filaments surrounded by a ring containing the adhesion-related cytoskeletal molecules, vinculin and talin. The space between the cells and the bone matrix constitutes the bone-resorbing compartment. Osteoclasts synthesize several proteolytic enzymes, such as cathepsin K and matrix metalloproteinase 9, which are transported toward the apical side of the cells and secreted into this compartment and play important roles in bone matrix degradation. Simultaneously, osteoclasts acidify this compartment by extruding protons via proton pumps (vacuolar type H-ATPase) located on the apical membrane, and the low pH (pH 3–4) of the compartment contributes to demineralization of the bone matrix. The apical membrane of the cells, so-called ruffled border membrane, develops a characteristic structure with numerous folds. The extensive folding of this structure is probably

Figure 1 A schematic representation of typical structures and functions of osteoclasts.

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due to the intense vesicular traffic associated with secretion as well as to the need for increasing the number of proton pumps via an amplification of the apical membrane. Osteoclasts contain high levels of tartrate-resistant acid phosphatase (TRAP), which is widely used as a specific histochemical marker of the cells. Calcitonin acts directly on osteoclasts and inhibits bone-resorbing activity of the cells, and abundant calcitonin receptors are expressed on the plasma membrane of mammalian osteoclasts, while specific calcitonin binding can hardly be observed in avian osteoclasts. Thus, osteoclasts are morphologically and functionally specialized for bone resorption [4]. Bone resorption by osteoclasts is necessary for normal skeletal development, for its adaptability, and for its maintenance. This process is critical for the growth, modeling, and remodeling of skeletal tissues, and under normal conditions it is tightly coupled to bone formation by osteoblasts (coupling). The disruption of this coupling process between bone resorption and formation leads to abnormal bone resorption under pathological conditions. III INVOLVEMENT OF OSTEOCLASTS IN BONE DESTRUCTION IN RA Rheumatoid arthritis is characterized by proliferative pannus formation leading to erosive bone destruction originating from the interface of cartilage and bone (the bare area). Synovial tissues of RA joints produce various kinds of inflammatory cytokines, such as interleukin-1 (IL-1) and tumor necrosis factor- (TNF-), which are believed to play important roles in joint destruction in the disease. The cellular mechanism underlining bone and cartilage destruction in RA still remains unclear, however, but recent studies have revealed the essential role of osteoclasts in bone and joint destruction in RA. Bromley and Wooley [5] observed a number of acid phosphatase–positive multinucleated cells (chondroclasts and osteoclasts) in the erosive areas of RA joints obtained at the time of joint replacements. In collagen-induced arthritis, multinucleated giant cells were observed at the bone–pannus junctions of arthritic joints, and cells isolated from the lesions can differentiate into mature osteoclasts [6]. Gravallese et al. also found multinucleated cells present on subchondral bone surface and in the areas of direct invasion of pannus into subchondral bone [7]. Their important discovery was that those multinucleated cells were positive for unique markers of osteoclasts such as TRAP, cathepsin K, and calcitonin receptors, satisfying the major criteria of mature osteoclasts. Interestingly, some multinucleated cells and mononuclear cells apart from the bone surface were TRAP-positive. These findings suggest the possible role of synovial tissues for osteoclastogenesis in RA. To reveal the osteoclastogenic potentiality of RA synovial tissues, synovial cells from RA synovia were cultured in the presence of osteotropic factors such as 1,25-dihydroxyvitamin D3 [1,25(OH)2D3] and macrophage colony stimulating factor (MCSF) [8]. After 3 weeks of culture, we observed many multinucleated giant cells, which were TRAPpositive, possessed abundant calcitonin receptors, and made resorption pits on dentine slices. We also demonstrated that peripheral monocytes can differentiate into osteoclastlike cells when cocultured with synovial fibroblasts obtained from RA synovial tissues in the presence of 1,25(OH)2D3 and MCSF. Similar results were reported by Fujikawa et al. [9]. They found that synovial macrophages isolated from RA synovial tissues can differentiate into osteoclast-like cells when cocultured with UMR106 rat osteoblast-like cells. These results suggest that RA synovial fibroblasts can support osteoclast differentiation from monocyte–macrophage lineage precursor cells under a suitable condition at least in vitro.

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IV ROLE OF RANKL/RANK PATHWAYS IN BONE DESTRUCTION IN RA Remarkable progress has been made during the last several years in the field of osteoclast research, primarily due to the finding of the receptor activator of NF-B ligand (RANKL)/RANK system [2]. RANKL is a member of the tumor necrosis factor superfamily cytokines, which was originally identified as a membrane-bound survival factor for dendritic cells produced by activated T cells [10]. The expression of RANKL can be also induced in osteoblasts and bone marrow stromal cells by osteotropic hormones such as 1,25(OH)2D3 and parathyroid hormone [11]. In the presence of MCSF, RANKL can stimulate osteoclast differentiation from hematopoietic precursor cells in vitro [11]. RANKL also acts on mature osteoclasts and activates the bone-resorbing activity and survival of the cells [12–14]. RANKL binds to its receptor RANK, a transmembrane receptor belonging to the TNF receptor superfamily, which is expressed in monocyte–macrophage lineage osteoclast precursor cells as well as in mature osteoclasts and dendritic cells [10,15,16]. Binding of RANKL to RANK induces intracellular signals including NF-B activation and cJun N-terminus kinase (JNK) activation. The other important actor in this system is osteoprotegerin (OPG), a soluble receptor of RANKL belonging to the TNF receptor superfamily [17]. Osteoprotegerin specifically binds to RANKL and inhibits RANKL activity by preventing its binding to RANK [2]. The essential role of RANKL/RANK signaling pathways in osteoclast development in vivo has been established by a series of targeted gene disruption experiments [18]. In short, the targeted disruption of either RANKL or RANK induced osteopetrosis in mice, a pathological bone disease which is characterized by an increased bone mass due to a deficiency in osteoclast differentiation. We and another group found that mice deficient in TRAF6, a signaling molecule found to be involved in RANK signaling, also showed osteopetrotic phenotypes [19,20]. In contrast, the targeted disruption of OPG induces reduced bone mass, reminiscent of osteoporosis, in mice due to the increased number and activity of osteoclasts [21,22]. These results clearly demonstrate the essential role of RANKL/RANK pathways in osteoclast development and activation in vivo. The next question is whether the RANKL/RANK system is also involved in pathological bone destruction such as RA. We and others have revealed that RANKL is highly expressed in synovial fibroblasts by Northern blotting, immunocytochemistry, and in situ hybridization [23–25]. 1,25(OH)2D3 treatment increased the expression of RANKL in synovial fibroblasts and reduced the expression of OPG in the cells [24]. RANKL expression was also detected in CD4 T lymphocytes in RA synovial tissues by in situ hybridyzation [23]. Kong et al. demonstrated that activated CD4 T lymphocytes fixed with paraformaldehyde or culture supernatants from activated T cells can support osteoclast differentiation through the surface-bound and/or soluble RANKL they produce [26]. They also showed that RANKL was expressed on the surface of activated T cells in synovial tissues of adjuvant arthritis rats [26]. These results suggest the important role of activated T lymphocytes in bone and joint destruction in RA. However, the role of T cells in osteoclast development is still controversial because activated T cells also produce many cytokines which inhibit osteoclast differentiation, such as interferon- and interleukin-10, as recently suggested by Takayanagi et al. [27]. In any case, these studies indicate that RANKL produced by synovial fibroblasts and/or activated T lymphocytes in RA synovial tissues may play an essential role in osteoclast development and bone destruction in RA. Based on these findings, Kong et al. proposed that OPG can be a potent

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Figure 2 Involvement of RANKL-RANK pathways in osteoclast differentiation and bone destruction in RA.

therapeutic agent against bone destruction in RA [26]. Exogenous administration of recombinant OPG suppressed bone and joint destruction in rat adjuvant arthritis. Interestingly, not only bone destruction but also cartilage destruction was prevented by OPG treatment [26]. Osteoprotegerin appears to be effective in other types of bone destruction as well. Honere et al. recently reported that OPG treatment blocks cancer-induced bone destruction and pain-related neurochemical changes in the spinal cord [28]. Clinical studies exploring the effect of OPG treatment for bone and joint destruction in osteoporosis, RA, and tumor metastasis are currently being conducted in the United States.

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V EFFICIENCY OF ADENOVIRUS VECTORS IN TRANSDUCING OSTEOCLASTS As mentioned, the important role of osteoclasts in bone and joint destruction in RA is now widely recognized, and pharmacological agents regulating osteoclast differentiation and/or function, such as OPG and bisphosphonates, can be potent therapeutic agents for RA. One alternative is gene therapy, where genes or cDNAs are directly transferred to target cells. In preclinical studies, ex vivo and in vivo gene transfer methods have been used successfully to reduce the joint destruction in experimental arthritis, and the first clinical trial, in which the IL-1 receptor antagonist gene was delivered to synoviocytes ex vivo, was started in 1996 in the United States [29,30]. Although modulating osteoclast function by gene therapy can be a good therapeutic approach for treating RA, transducing foreign genes into osteoclasts even in vitro is extremely difficult mainly because osteoclasts are terminally differentiated nonproliferating cells with very short life span. Antisense technology has been successfully utilized to suppress particular genes’ expression in osteoclasts [31–33]. We previously reported that osteoclast function was successfully suppressed by inhibiting the expression of c-src or c-cbl by antisense oligodeoxynucleotides [33]. The shortcoming of antisense technique is that the antisense inhibition does not always work, and that it does not enable us to “overexpress” any genes. Moreover, efficient delivery of antisense oligonucleotides to the target cells in vivo is technically demanding [34]. Retrovirus vectors can be used to transduce osteoclast precursors, and Matsuo et al. reported that the retrovirus vector–mediated c-fos gene or fra-1 gene could rescue osteoclast differentiation from c-fos–deficient precursor cells [35]. Retrovirus vectors can transfect by integrating the transgene into target cell genome, which leads to more prolonged gene expression. However, retrovirus vectors penetrate the nucleus only at mitosis, so transfection is restricted to proliferating cells, and therefore they do not transduce postmitotic cells such as mature osteoclasts. In addition, retrovirus vectors usually give inefficient gene transfer for most cell

Figure 3 Effective gene transduction into mature osteoclasts by adenovirus vectors. Human osteoclast-like cells obtained from giant cell tumors were infected with either (A) control virus or (B) LacZ virus, and stained for -galactosidase activity. Many multinucleated cells infected with LacZ virus were positively stained, indicating an efficient gene transduction. Bar  100 m. (From Ref. 40.)

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types in vivo [36]. It is widely recognized that adenovirus vectors give quite efficient gene transfer to many cell types, especially nonproliferating quiescent cells [37,38]. We recently found that adenovirus vectors can be efficiently utilized as gene transfer agents for posmitotic osteoclasts both in vitro and in vivo [39,40]. Adenovirus vectors have several advantages in introducing foreign genes into mature osteoclasts. First, these vectors are capable of infecting a variety of terminally differentiated cells, such as neurons and hepatocytes. Second, recombinant adenoviruses can be easily amplified to a very high titer in vitro. Third, adenovirus infection to the cells has been reported to require the interaction of the RGD sequence in the penton base of the virus with the cell surface vitronectin receptors (v3 or v5 integrins), and v3 integrin is expressed at very high levels on the cell surface of osteoclasts. Recombinant adenovirus carrying the lacZ gene can infect human osteoclast-like cells (OCLs) obtained from giant cell tumors as well as mouse OCLs formed in vitro (Fig. 3). At an MOI of 100, more than 85% of OCLs were positively stained by galactosidase (-gal) activity with no apparent morphological changes or cellular toxicity. When injected in bone marrow cavity of adult mice femur, many osteoclasts present on bone surface were positively stained for -gal activity [40]. Our results suggest that recombinant adenovirus vector system is suitable for gene transduction into osteoclasts. VI MODULATION OF OSTEOCLAST FUNCTION BY REGULATING C-SRC ACTIVITY c-src was first identified as the normal cellular counterpart of the transforming protein encoded by Rous sarcoma virus, v-Src [41,42]. The protooncogene product c-Src is a 60-kD protein and belongs to nonreceptor-type tyrosine kinase family, i.e., Src family tyrosine kinase family. The c-src protooncogene is highly conserved throughout evolution and widely expressed. Although the level of the expression is low in most cell types, some cell types, particularly neurons and blood platelets, express high levels of c-Src protein. It is known that c-Src and the other members of the Src family, which share highly conserved sequences both within and outside the kinase catalytic domain, play important roles in signal transduction mechanisms that contribute to the regulation of cell growth and development [41,42]. However, the physiological role of the c-src gene had not been clarified until Soriano et al. successfully performed the targeted disruption of the gene by homologous recombination in mouse embryos in 1991 [43]. Unexpectedly, cell proliferation and other basal functions did not appear to be impaired in c-src–deficient animals, and no obvious phenotypic or functional abnormalities were observed in neuronal tissues or in platelets, probably because other Src family members are expressed in these cells, which might impart a degree of functional redundancy with c-src [43]. Surprisingly, the mice showed striking skeletal abnormalities with a phenotype of osteopetrosis. These included a failure of the incisors to erupt, a slower growth, shorter and abnormally shaped long bones, and decreased bone marrow cavities, which were all due to reduced bone resorption. In vitro osteoclast formation experiments and in vivo bone marrow transplantation studies have revealed that osteoclast differentiation was not impaired, but that bone-resorbing activity of mature osteoclasts was much reduced in c-src knock-out (KO) mice [44,45]. The morphological characteristics of the KO mouse osteoclasts was their disorganized ruffled border structure. The ruffled border is the apical membrane of the osteoclast, which is extensively folded due to the intense vesicular traffic associated with proton and lysosomal enzyme secretion. We and others demonstrated that c-Src is highly expressed in osteoclasts, and c-Src protein is highly concentrated on ruffled border membranes and intracellular membranes

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(Fig. 1) [46,47]. This, in addition to the association of this molecule with secretory granules and vesicles in platelets, chromaffin cells, and neurons, leads us to consider that c-Src may contribute vesicle targeting or membrane fusion in osteoclasts and plays essential roles for osteoclastic bone resorption. The fact that no other abnormalities in c-src KO mice were found outside the skeletal tissues leads us to consider that c-Src can be an ideal therapeutic target for suppressing pathological bone resorption by inhibiting osteoclast function without affection other tissues or cells. In fact, several specific inhibitors to c-Src kinase have been developed and reported to be effective in suppressing animal bone resorption models [48,49]. The tyrosine kinase activity of c-Src is strictly regulated by phosphorylation and dephosphorylation of the tyrosine residue located close to the C-terminus, which corresponds to tyrosine 527 (Tyr527) in chicken c-Src [41]. Dephosphorylation of this residue causes a 10- to 20-fold increase in the kinase activity of c-Src. C-terminus Src family kinase (Csk) is a cytoplasmic tyrosine kinase which specifically phosphorylates Tyr527 of c-Src and thereby negatively regulates its kinase activity [50]. In order to modulate osteoclast function by regulating c-Src kinase activity, we constructed adenovirus vectors encoding wild type csk gene (Csk virus) and kinase-deficient mutant of the gene (Csk-KD virus). These viruses efficiently induced the overexpression of the genes in osteoclasts. Csk virus dosedependently inhibited the kinase activity of c-Src in osteoclasts, whereas Csk-KD virus infection rather increased the kinase activity, probably by working as a dominant negative molecule against intrinsic Csk or Csk-related molecule in osteoclasts. Adenovirus vector–mediated Csk overexpression induced dramatic cytoskeletal disorganization in osteoclasts and strongly inhibited pit formation on dentine slices. To examine the effect of these adenoviruses on in vivo bone resorption, we utilized a modified bone resorption model reported by Uy et al. [51]. Either Csk virus or Csk-KD virus was injected onto the calvaria of adult mice together with interleukin-1 (IL-1) twice a day for 3 days. The mice were then sacrificed and the calvaria were examined morphologically. Interleukin-1 injection induced a significant increase in the area of bone resorption and in the number of osteoclasts, and Csk virus injection significantly reduced the bone resorption induced by IL-1 (Fig. 4), while Csk-KD virus injection rather exacerbated it (not shown). This suggests that Csk virus efficiently suppresses bone-resorbing activity of osteoclasts not only in vitro but also in vivo. Interestingly, periosteal inflammatory reaction caused by IL-1 injection was also reduced by Csk virus. Because Src kinase is known to be involved in the cell cycle progression in fibroblasts and required for induction of DNA synthesis of various growth factors such as platelet-derived growth factor [42], we suspect that Csk virus also has an inhibitory effect on the proliferation of periosteal fibroblastic cells. VII AMERIOLATION OF ARTHRITIC BONE DESTRUCTION BY ADENOVIRUS VECTOR–INDUCED CSK OVEREXPRESSION Although the exact etiology of RA remains unclear, previous studies have revealed the involvement of various types of cells such as T and B lymphocytes, macrophages and synovial fibroblasts, and cytokines and growth factors they produce. Therefore, the ideal therapy of RA cannot be accomplished only by targeting a single type of cell or a particular molecule in spite of the remarkable clinical success of the anticytokine therapies. Src family members of tyrosine kinases are involved in signal transduction pathways that regulate a variety of biological activities [42]. In T and B cells, they are activated in response to immune recognition receptor stimulation as well as cytokine stimulation. C-Src is found to be

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Figure 4 Calvarial sections of mice killed on day 3 of IL-1 injection (A). Significant increase in bone resorption was present along the marrow–bone interface in bone from a mouse treated with IL-1. (B) The bone resorption, periosteal cell proliferation, and inflammatory reaction are less pronounced in the calvaria of the mouse treated with IL-1 and Csk virus than in the IL-1–treated mouse. (From Ref. 55.)

involved in the cell cycle progression and growth factor signaling in fibroblasts. Src kinases are also implicated in adhesion events regulated by integrins, cadherins, selectins, and CAMs. These results suggest that Src family kinases are implicated in various aspects of cellular events in the pathogenesis of RA, and regulating Src activity can be a promising approach for the treatment of RA. As previously reported, adenovirus vectors are quite effective in delivering genes to synovial fibroblasts both in vitro and in vivo [52]. The effect of modulation of Src kinase activity in synovial fibroblasts was investigated by utilizing adenovirus vector–mediated Csk overexpression. Synovial fibroblasts were obtained from synovial tissues of

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RA patients who fulfilled the American College of Rheumatology criteria of RA. As shown in Fig. 5, Csk overexpression dose-dependently suppressed the proliferation of the cells without causing any toxic effects. The expression of interleukin-6, one of the major inflammatory cytokines produced by synovial fibroblasts, was also reduced by Csk virus expression as revealed by Northern blot analysis. These results suggest that Csk virus can inhibit hyperplasia of synovial fibroblasts as well as osteoclastic bone resorption. We next examined the effect of Csk virus on pathological bone resorption in an animal model of RA. There are two types of gene transduction strategies for RA; systemic administration and local administration. The efficient in vivo gene delivery to synovial cells by local administration of adenovirus vectors has been well established [52]. As shown in Fig. 5, strong -gal staining of synovial lining cells was observed by intraarticular injection of the adenovirus vector encoding lacZ gene as previously reported. In addition, TRAP-positive osteoclasts along the erosive bone surface demonstrated strong -gal staining as shown in the serial tissue sections, indicating that intraarticular injection of adenovirus vectors can transduce osteoclasts in vivo. Western blot analysis demonstrated that Csk virus injection led to a 12-fold increase in the expression of Csk protein in the injected joints on day 7, and the expression returned to the control level on day 42. The effect of Csk adenovirus on inflammatory joints in adjuvant arthritis rats was investigated by direct administration of the virus on their ankle joints. Not only was the bone destruction by osteoclasts suppressed by Csk virus injection, but also synovial inflamma-

Figure 5 Effect of Csk overexpression on proliferation and cytokine production of RA synovial fibroblasts. Synovial fibroblasts obtained from RA patients were infected with (A) control virus or (B) Csk virus, and cell proliferation was assayed by determining bromodeoxyuridine (BrDU) incorporation into the nuclei of the cells. (A) Most nuclei of the cells were positively stained (black staining) by anti-BrDU antibody. (B) Positive staining was hardly observed in the nuclei of Csk virus– infected cells. (C) Csk overexpression strongly suppressed IL-6 expression in the cells by Northern blot analysis. (From Ref. 39.)

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Figure 6 Adenovirus-mediated gene transduction in osteoclasts in vivo. LacZ virus was injected into the inflammatory ankle joint of an adjuvant arthritis rat, and the expression lacZ gene in osteoclasts was determined in the serial sections by enzyme histochemistry of (A) -galactosidase and (B) TRAP after 1 week of the viral injection. Most of the TRAP-positive osteoclasts were positively stained for -galactosidase activity. (From Ref. 39.)

tory reaction detected by arthritis score or paw swelling was reduced (Figs. 7 and 8). The mechanism of this anti-inflammatory effect of the virus is probably due to its suppressive effect on the proliferation of synovial fibroblasts and their production of inflammatory cytokines such as interleukin-6 as we reported.

Figure 7 X-ray analysis of the effect of Csk virus on the joint destruction of adjuvant arthritis rats. Either (A) control virus or (B) Csk virus was injected into the inflammatory ankle joints of adjuvant arthritis rats after 7 days of adjuvant injection. Bone destruction and periarticular osteoporosis was reduced in Csk virus–injected animals.

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Figure 8 Histological analysis of the effect of Csk adenovirus injection on adjuvant arthritis rats. (A) Proliferating synovial cells are invading into the calcaneus of the control virus–injected rat (arrowhead). (B) bone erosion was significantly reduced in the Csk virus–injected animal.

VIII CONCLUDING REMARKS The ultimate goal of the treatment of RA is to prevent the bone and joint destruction and preserve the daily activity of the patients. Recent studies have revealed that osteoclasts are involved in the pathogenesis of the bone and joint destruction in RA and can be a potent therapeutic target of the disease [7,8,23–25]. Therapies that inhibit osteoclast formation or function can at least ameliorate the progression of these bone changes. Kong et al. clearly demonstrated that blocking RANKL/RANK pathways by systemic administration of OPG could suppress not only bone destruction, but also cartilage destruction in adjuvant arthritis in rats [26], and several other groups proposed bisphosphonates, potent antiresorption agents, as possible therapeutic agents for arthritic joint destruction [53,54]. However, inhibition of osteoclast function by bisphosphonates or calcitonin alone could not completely prevent bone erosion in RA in spite of their preventive effects against systemic bone loss [53]. Therefore, the combination of anti–bone resorption therapy and anti-inflammatory therapy can be an ideal therapy of RA. In this regard, suppressing Src family kinase activity can be a good therapeutic approach to RA because Src family members of tyrosine kinases are implicated in many signal transduction pathways that affect various aspects of RA pathology. They are involved in the mitogenic response to several growth factors in fibroblasts; activation of lymphocytes, platelets, and osteoclasts; and various cytokine signaling. As previously mentioned, adenovirus vector–mediated csk gene expression can be a promising means of preventing arthritic bone destruction by suppressing osteoclast function as well as activation of synovial fibroblasts. Of course we have to realize the disadvantages as well as advantages of using adenovirus vectors as therapeutic reagents [36]. The disadvantages of the adenovirus vectors include the transient gene expression because they do not integrate the transgene into target cell chromosome, immunological reaction such as neutralizing antibody response and cytotoxic T lymphocyte responses against the virus, and the dissemination of the vectors from the site of local injection. The safety issue is particularly important in case of its clinical application, and in fact the first case of fatality induced by the infusion of adenovirus vectors into hepatic artery was recently reported. Therefore, development of a new generation of adenovirus vectors is absolutely necessary for the clinical usage of the vector to RA gene therapy.

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49. Yoneda T., Lowe C., Lee C. H., Gutierrez G., Niewolna M., Williams P. J., Izbicka E., Uehara Y., Mundy G. R. 1993. Herbimycin A, a pp60c-Src tyrosine kinase inhibitor, inhibits osteoclastic bone resorption in vitro and hypercalcemia in vivo. J. Clin. Invest. 91:2791–2795. 50. Okada M., Nada S., Yamanashi Y., Yamamoto T., Nakagawa H. 1991. CSK: a protein–tyrosine kinase involved in regulation of src family kinases. J. Biol. Chem. 266:24249–24252. 51. Uy H. L., Dallas M., Calland J. W., Boyce B. F., Mundy G. R., Roodman G. D. 1995. Use of an in vivo model to determine the effects of interleukin-1 on cells at different stages in the osteoclast lineage. J. Bone Miner. Res. 10:295–301. 52. Nita I., Ghivizzani S. C., Galea-Lauri J., Bandara G., Georgescu H. I., Robbins P. D., Evans C. H. 1996. Direct gene delivery to synovium. An evaluation of potential vectors in vitro and in vivo. Arthritis Rheum. 39:820–828. 53. Eggelmeijer F., Papapoulos S. E., van Paassen H. C., Dijkmans B. A., Valkema R., Westedt M. L., Landman J. O., Pauwels E. K., Breedveld F. C. 1996. Increased bone mass with pamidronate treatment in rheumatoid arthritis. Results of a three-year randomized, double-blind trial. Arthritis Rheum. 39:396–402. 54. Zhao H., Shuto T., Hirata G., Iwamoto Y. 2000. Aminobisphosphonate (YM175) inhibits bone destruction in rat adjuvant arthritis (in press). J. Orthop. Sci. 5:397–403. 55. Miyazaki T., Takayanagi H., Isshiki M., Takahashi T., Okada M., Fukui Y., Oda H., Nakamura K., Hirai H., Kurokawa T., Tanaka S. 2000. In vitro and in vivo suppression of osteoclast function by adenovirus vector–induced csk gene. J. Bone Miner. Res. 15:41–51.

24 Muscle-Derived Cell–Based Gene Therapy and Tissue Engineering for the Musculoskeletal System Nobuo Adachi, Dalip Pelinkovic, Kenji Sato, Freddie H. Fu, and Johnny Huard Children’s Hospital of Pittsburgh and University of Pittsburgh Musculoskeletal Research Center, Pittsburgh, Pennsylvania

I INTRODUCTION Treatment of musculoskeletal disorders has improved through the advances in biological and biomechanical research in the past two decades. However, despite considerable progress, no definitive treatment has been established as a gold standard for musculoskeletal structures that have a low intrinsic healing capacity, such as the cruciate ligament, the meniscus, and cartilage. Even tissues that have a good healing capacity, including muscle and bone, sometimes show incomplete healing, causing a considerable decrease of patient quality of life or requiring repeated treatments. Various growth factors have been identified to enhance healing in the tissues of the musculoskeletal system. Although some studies have revealed that direct injection of therapeutic proteins has beneficial effects on tissue healing, their relatively short half-lives in vivo often make a high dosage of protein or repeated injection necessary. Moreover, there is a possibility that unregulated high dosage or repeated injection of growth factors can influence not only the injured tissues, but also normal structures and can lead to adverse effects. Gene therapy represents a promising method to deliver efficient and adequate amounts of growth factors into the injured tissues for an extended period of time. This technique relies on the alteration of the cellular genetic information of the target cells. In orthopedic fields, defined genes encoding growth factors are transferred into the target tissue using viral or nonviral vectors through different gene transfer strategies, including local and systematic delivery. 483

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The other important approach for the repair of the musculoskeletal system is tissue engineering. Tissue engineering is a technology based on the development of biological substitutes for the repair, reconstruction, regeneration, or replacement of biological tissues. Recently, genetic modification has been included in tissue engineering to optimize the treatment of various tissues. The combination of gene therapy and tissue engineering could help to develop new approaches to improve the healing of various tissues of the musculoskeletal system that display low healing capacities. Tissue engineering for the musculoskeletal system requires three major components: a cell source, a scaffold, and growth factors. Until now, various cells—from undifferentiated pluripotent cells, such as embryonic stem cells, to welldifferentiated cells, such as chondrocytes—were used as cell sources for tissue repair in orthopedic fields. Our research team has focused its work on genetically modified musclederived cells (MDCs) for muscle-related disorders such as Duchenne muscular dystrophy and sports-related muscle injuries. Recently, the MDCs have been studied and applied to other musculoskeletal tissues, including bone, ligament, meniscus, and cartilage. This chapter provides an overview of the current status of our research on gene therapy and tissue engineering for musculoskeletal disorders. We will also discuss various hurdles that have to be circumvented for general application of this technology to the musculoskeletal system and also future directions. II MUSCLE INJURY AND REPAIR Muscle injuries (contusions, lacerations, and strains) are by far the most common injuries in sports, their incidence varying from 10 to 55% of all injuries sustained in sporting events [1,2]. Muscle injuries may result from both direct (contusions, lacerations) and indirect trauma (strains, ischemia, and neurological injuries). The majority of muscle injuries (90%) are caused either by contusion or by excessive strain of the muscle. Skeletal muscle can regenerate extensively after injuries, including contusions, lacerations, and strains [3–10]. The myogenic cells responsible for this regeneration are the mononucleated satellite cells located between the basal lamina and plasma membrane of the muscle fiber [11,12]. It is hypothesized that after muscle injury, disruption of the basal lamina and plasma membrane releases and activates the satellite cells [11,12]. The satellite cells begin to proliferate and differentiate into multinucleated myotubes and eventually into myofibers. The growth of these regenerating myofibers at the injured site promotes the healing of the muscles, although very slowly and occasionally with an incomplete functional recovery. The events in muscle tissue repair have been observed in both animal models and clinical experience; however, muscle injuries in medical practice are usually found to be more destructive than those in animal models [13]. The injured gap is usually filled with hematoma, which matures into proliferating granulation tissue and eventually results in the formation of connective scar tissue. These events not only make the repair complex, but may also inhibit the complete regeneration of muscle tissue and contribute to the partial functional recovery [13]. Characterization of approaches to improve muscle healing after injury requires the development of well-defined and reproducible animal models of muscle injury. Animal models of muscle contusion, laceration, and strain have been previously reported. Our research team has also established reproducible orthopedic muscle injury models (contusion, laceration, and strain) in mice [3–7]. Under these conditions, muscle myofiber regeneration is found at 7 and 10 days after injury, but begins to decrease at 14 days and continues decreasing until 35 days. Concomitantly, fibrosis is developed beginning at 14 days and grad-

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ually increases until 35 days. Fibrosis appears at the time of muscle regeneration and, therefore, hinders the healing response. These results show that muscle healing after injuries is still incomplete at 1 month postinjury, and approaches that can both enhance muscle regeneration and prevent muscle fibrosis may lead to improved muscle healing after injuries. A Growth Factors and Muscle in Vitro Injured skeletal muscle releases numerous growth factors acting in autocrine and paracrine fashion to modulate muscle healing. These proteins activate satellite cells to proliferate and differentiate into myofibers. The delivery of exogenous growth factors specifically selected to enhance myofiber regeneration is an intuitive therapeutic approach to muscle injuries. In vitro experiments have identified several growth factors that can enhance myogenic proliferation and differentiation. Myoblast proliferation and differentiation was assessed at 48 and 96 h after incubation in selected growth factors. Insulin-like growth factor-1 (IGF-1), basic fibroblast growth factor (bFGF), and nerve growth factor (NGF) significantly enhanced myoblast proliferation; and IGF-1, bFGF, NGF, and acidic fibroblast growth factor (aFGF) increased myoblast differentiation into myotubes. Consequently, IGF-1, bFGF, and NGF are the logical candidates for therapeutic applications to enhance muscle healing [3–7]. B The Effects of Growth Factors in Injured Skeletal Muscle in Vivo The technique chosen to deliver specific growth factors to injured muscles is of paramount importance to optimize therapeutic benefit. Three different approaches have been used and tested to deliver the growth factors into injured muscle: direct injection of the growth factor protein, gene therapy based on direct and ex vivo gene transfer of viral vectors, and, finally, myoblast transplantation. 1 Direct Injection of the Growth Factor Protein We injected IGF-1, bFGF, and NGF into injured muscles (contusions, lacerations, and strains). Muscle healing was analyzed by histology and physiological testing (contractile properties). We observed that individual direct injection of IGF-1, bFGF, and NGF into injured muscles can increase the number of regenerating myofibers and improve fast twitch and tetanic muscle strength 15 days after injury. Of those growth factors, IGF-1 had the best overall effect on the healing process of the injured muscle by increasing both muscle regeneration and muscle strength [7]. 2 Gene Therapy Based on Viral Vectors Direct injection of growth factors is hindered by the high amount of factor required to have a significant effect as well as by the bloodstream’s rapid clearance of the substances. If a dose-dependent response resulting in better muscle healing occurs with higher doses of growth factors, gene therapy may offer new opportunities because a high level and stable expression of these substances in the injured muscle may be achieved. In fact, we found that growth factors exhibited a dose-dependent effect in vitro on the myoblast proliferation and differentiation, while in vivo three consecutive injections of a relatively high concentration of growth factor (100 ng) were required to achieve enhancement of muscle healing [3,5,7]. Direct gene therapy to deliver genes to skeletal muscle is possible using naked DNA, adenovirus, retrovirus, herpes simplex virus, and adeno-associated virus. Most of

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these vectors transduce relatively few mature myofibers. However, adenovirus can transduce a large number of regenerating muscle fibers, a condition present in injured muscle. We have investigated the ability of adenoviral vectors to achieve gene transfer to the injured muscle. Direct intramuscular or myoblast-mediated ex vivo gene transfer of recombinant adenovirus carrying the -galactosidase reporter gene was highly capable of transducing injured muscle [3,5]. Many LacZ-expressing myofibers have been identified in the injured sites of contused, lacerated, and strained muscles at 5 days after either direct or ex vivo gene transfer. Because IGF-1 was by far the most potent growth factor to stimulate muscle regeneration, we engineered an adenovirus that carries the gene encoding for IGF-1. The viral transduction of myoblasts in vitro led to nanogram quantities of IGF-1 secreted by the transduced cells [14]. However, after direct injection of the adenovirus IGF-1 (Ad-IGF-1) vector in the injured muscles, no significant improvement of muscle strength (fast twitch and tetanic strength) was found when compared with the control muscles, which were injured and then injected with saline. However, the myoblast-mediated ex vivo gene transfer by adenovirus IGF-1 (myob/ad-IGF-1) did improve muscle healing after laceration in immunocompetent mice, but at a level lower than that observed with the transplantation of the same number of myoblasts alone (without IGF-1). Although immune response against the adenovirus may limit the success of gene therapy, the lack of improvement in severe combined immune deficient (SCID) mice suggests that the immune response is not a major factor in this lack of improvement of muscle healing. The histology of the injected SCID mice’s muscle, which was lacerated and injected with adenovirus IGF-1 and IGF-1–expressing myoblasts, showed the development of muscle fibrosis within the lacerated site. These results taken together strongly argue that a high level of IGF-1 mediated by adenoviral-based gene therapy will block muscle fibrosis and significantly improve muscle healing after injury. 3 Autologous Myoblast Transfer Myoblast transplantation has been extensively studied for increasing the muscle mass in dystrophic muscle, such as in Duchenne muscular dystrophy (DMD). This approach has been found capable of enhancing muscle regeneration, leading to an improvement of muscle mass and muscle strength in dystrophic muscle [15–26]. We have also observed that the injection of adenovirally transduced primary myoblasts in notexin-injured muscle (snake venom) enhances muscle regeneration [24–26]. In fact, numerous transduced myofibers have been observed, showing that myoblast transplantation can enhance muscle regeneration in the injured muscle [24–26]. To prevent immune complications, we recently injected autologous myoblasts in a mouse laceration model. We observed that the injection of normal myoblasts in lacerated muscle could physiologically improve the strength of the muscle in normal immunocompetent mice when compared with the control injured muscle injected with saline. As observed with direct injection of IGF-1, the implantation of myoblasts enhances muscle regeneration and improves muscle healing after injury, but the full recovery of the injured skeletal muscle is still hindered by the development of fibrosis. III DEVELOPING AN APPROACH TO PREVENT FIBROSIS AND IMPROVE MUSCLE HEALING Large formations of scar tissue in injured muscles at 1 month postinjury suggest that the natural healing of contused, lacerated, and strained muscles remains incomplete. Fibrosis

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has been found to appear at the second week after muscle injury and enlarges over time. This significantly impairs muscle regeneration [2–7]. We have also observed that the injection of IGF-1 protein, adenovirus that encodes for IGF-1, and myoblasts improves muscle healing, but the development of fibrosis greatly limits the full recovery of the injured skeletal muscle. It is therefore likely that approaches to minimize the development of fibrosis after injury may lead to a full recovery of the injured muscle. We investigated the use of antifibrosis agents to improve muscle healing after injury. A Use of an Antifibrosis Agent to Improve Muscle Recovery We subsequently investigated the use of decorin, a proteoglycan that acts as an inhibitor of TGF- [27–33], one of the major causes of tissue fibrosis [34–36], as an approach to prevent muscle fibrosis. The direct injection of human recombinant decorin into the lacerated muscle in vivo decreased the amount of fibrosis in the injured muscle [37]. In detail, 5, 25, and 50 g of decorin were injected into the lacerated muscle at 2 weeks postinjury. Levels of muscle regeneration and fibrosis were monitored at 14 days post–decorin administration. The area of muscle fibrosis, as measured by histology and immunofluorescence for vimentin, decreased in the injured muscle treated with decorin when compared with the lacerated control muscle injected with saline. Compared with the control muscle, the injured site treated with decorin had higher numbers of regenerating myofibers with larger diameters. Finally, we observed that decorin had a beneficial effect on the functional recovery of lacerated muscle. Indeed, we observed that the injection of 50 g of decorin at 15 days postlaceration significantly improved (p  .001) both the fast-twitch and tetanic strength of the lacerated muscle when compared with the lacerated control injected with saline. More importantly, the strength of the injured muscle treated with decorin was found to be similar (no significant difference) to that observed in a normal, noninjured muscle. The injection of decorin improved both the structure and function of the lacerated muscle to a near complete recovery. As for the growth factors, the successful clinical implementation of this technique is currently limited by the problem of maintaining an adequate concentration of therapeutic protein in the lesion site or target tissue. The short half-lives of growth factors and decorin, as well as systemic lavage, may lead to a rapid clearance of the substances from the desired site. To address these issues, gene therapy may be an interesting system of delivery to the muscle. A large number of recreational and professional athletic injuries involve skeletal muscle. By enhancing muscle regeneration and preventing fibrosis, it may be possible to promote efficient muscle healing and complete functional recovery and, moreover, to reduce the risk of reinjury after sports-related muscle injuries. Further studies are investigating the dose-dependent response; the optimal time period for injection; the potential synergetic effect of the association of growth factor and decorin; long-term follow-up; and the application of this treatment to different muscle conditions, such as contusion, strain, and ischemia. The results of such investigations may provide ideal treatments for these common muscle injuries. IV BONE HEALING Although the majority of fractures do heal well, delayed unions or nonunions sometimes occur, and their management can be devastating. Orthopedic surgeons also have a chance

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to treat large bone defects after tumor resection or trauma. Beyond fracture repair and bone defects, bone formation and healing have major implications in other orthopedic fields, including implant ingrowth, avascular necrosis, and spinal fusion. Bone healing is an important matter not only for orthopedic surgery, but also for other clinical fields, such as plastic and maxillofacial surgery. We usually treat bone defects or nonunions of the fracture with bone autografts, vascularized bone grafts, allografts, bone transports, or electrical stimulation. However, patients usually endure long recovery periods with numerous procedures, potential donor morbidity, and often less than satisfactory results. Therefore, research toward improving bone healing through mechanical and biological approaches is an area of intense interest. We investigated the feasibility of using genetically modified MDCs for bone defects. In 1999, Day et al. created segmental bone defects in rabbit tibias and injected MDCs that had been obtained and transduced with adenovirus encoding the LacZ gene [38]. Six days after injection, all of the rabbits injected with transduced MDCs demonstrated -galactosidase production in the defects and the muscles surrounding the defects macroscopically and histologically, indicating the feasibility of using MDCs to deliver genes for the treatment of bone defects. One important approach for bone healing is the use of growth factors that enhance bone healing. There are many growth factors that have osteoinductive and angiogenic potentials, such as bone morphogenetic protein (BMP), transforming growth factor-, and vascular endothelial growth factor (VEGF). Bone morphogenetic proteins are a family of growth factors originally described by Urist that regulate bone osteogenesis, bone healing, and ectopic bone formation [39]. Bone morphogenetic protein-2 has been well studied and used in research about bone healing. Recently, Bosch et al. reported MDC-mediated ex vivo gene therapy to induce bone formation using an adenoviral vector encoding BMP-2 (Ad-BMP-2) [40]. They transduced the mouse MDCs with Ad-BMP-2 and injected them into the triceps surae muscles of the hind limbs of severe combined immune deficient mice. Heterotopic bone formation in the muscle site injected with Ad-BMP-2–infected cells was evident by 2 weeks both radiographically and histologically. Nodules of bone continued to enlarge through 4 weeks and were grossly apparent at that time. The bone produced contained osteoid and bone marrow elements, as evidenced by hematoxylin and eosin staining and von Kossa staining for mineralization. In addition to the ex vivo approach, in 1999, Musgrave et al. reported adenovirus-mediated direct gene transfer of BMP-2 [41]. They injected the Ad-BMP-2 directly into the triceps surae of the SCID mice and normal adult immunocompetent mice. They confirmed ectopic bone formation macroscopically and histologically in the muscle injected with Ad-BMP-2 as early as 2 weeks in SCID mice and after 3 weeks in normal immunocompetent mice. However, the gross ectopic bone formation in normal mice was inferior to that observed in SCID mice. In 2000, Lee et al. reported the feasibility of using MDCs as a vehicle for therapeutic gene delivery to a skull defect model. Primary MDCs from normal male mice were transduced with adenovirus encoding the recombinant human BMP-2 gene [42]. These cells were implanted into nonhealing skull defects in female SCID mice using collagen sponges as a scaffold. They reported that skull defects treated with MDCs genetically engineered to express rhBMP-2 had significantly better closure of the defects at both 2 and 4 weeks after implantation. Fluorescent in situ hybridization (FISH) using a Y chromosome–specific probe showed that the majority of transplanted cells were located at the periphery of newly formed bone. Moreover, the implanted cells found within the new bone

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colocalized with osteocalcin staining, indicating that they differentiated in vivo into osteogenic cells. These data demonstrated the feasibility of MDC-based gene therapy for bone healing. This MDC-based gene therapy to produce bone can be applied to various skeletal disorders. Genetically modified MDCs capable of bone formation may be exploited to reconstruct the bone defects and minimize the use of autografts, allografts, and bone transportation.

V INTRA-ARTICULAR DISORDERS A Articular Cartilage Articular cartilage is a highly organized tissue composed of water, chondrocytes, and an extracellular matrix including collagens, proteoglycans, and other minor components [43,44]. The most characteristic feature of articular cartilage is its avascular status. Therefore, articular cartilage defects that do not penetrate subchondral bone cannot heal because cell infiltration from the bone marrow does not occur. On the other hand, full-thickness cartilage defects penetrating subchondral bone heal with fibrocartilaginous tissue, which is mainly composed of type I collagen. It has different biochemical and biomechanical properties than normal cartilage. This fibrocartilaginous tissue cannot resist mechanical stress for a long time and tends to deteriorate with time. These facts mean that if articular cartilage is injured, it cannot heal fully. More importantly, injury can lead to secondary arthritis, causing a considerable decrease of quality of life for the patient and tremendous health care costs. Numerous attempts to repair cartilage have been conducted in basic and clinical research, including abrasion arthroplasty, microfracture, and transplantation of chondrocytes [45,46], perichondrium, periosteum, meniscal allografts [47,48], and osteochondral grafts [49]. However, no treatment yet has regenerated long-lasting hyaline cartilage. In 1994, Brittberg et al. reported that cartilaginous defects in the knee were treated successfully with transplantation of chondrocytes cultured in monolayer [45]. This study was a breakthrough for the treatment of cartilage injuries, and it has been performed widely, especially in the United States and Europe. However, two concerns were raised about this chondrocyte transplantation using the monolayer culture system. One is the maintenance of chondrocyte phenotype during the prolonged monolayer culture period. The other is that there is a risk of transplanted chondrocyte leakage from the grafted site during range of motion and weight bearing of the joint. One solution for these concerns is using a three-dimensional culture system, such as collagen gels, fibrin glue, alginate, and agarose. It has repeatedly been demonstrated that chondrocytes can be cultured in a three-dimensional culture system without altering their phenotype and losing the ability to accumulate extracellular matrix composed of glycosaminoglycans and type II collagen, forming an architecture that resembles hyaline cartilage. Recently, Katsube et al. reported transplantation of chondrocytes cultured in type I collagen gels [46]. They showed better repair of the defects histologically with transplantation of chondrocytes cultured in collagen gels than with transplantation of chondrocytes cultured in monolayer, periosteal graft only, or no treatment. We believe that one of the best procedures for full-thickness cartilage defects is transplantation of chondrocytes embedded in collagen gels. Another important approach for cartilage repair is using growth factors that can enhance cartilage healing. Several growth factors, including TGF-1 [50,51], IGF-1 [52,53],

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BMP-2 [54,55], bFGF [56,57], and interleukin-1 receptor antagonist, have been found capable of enhancing chondrocyte proliferation and extracellular matrix synthesis in vitro and in vivo. As mentioned previously, when applying the growth factors for cartilage healing, their relatively short half-lives in vivo often necessitate a high dosage of protein or repeated injections. Unregulated high dosage or repeated injections can cause adverse effects, such as hypertrophy of cartilage or osteophyte formation. Gene therapy can be the best solution to these problems. It can deliver efficient amounts of growth factors into cartilage defects for an extended period of time. Chondrocyte transplantation is a good candidate for the ex vivo approach to cartilage repair because in its original technique, chondrocytes are harvested from the host before transplantation. This gives us a chance to modify the cells genetically before transplantation. With the ex vivo gene therapy approach, cells are isolated and cultured, and then the desired gene is inserted into the cells via viral or nonviral vectors. Then genetically modified cells are transplanted to the patient. Chondrocyte transplantation is a logical candidate for the ex vivo approach for cartilage repair. While offering advantages such as the safety of genetic manipulation, the ex vivo gene therapeutic approach is technically more challenging than direct gene therapy procedures. In 1997, Kang et al. used the ex vivo approach to transplant chondrocytes that had been transduced with retrovirus carrying the LacZ gene to the full-thickness cartilage defects created in the medial femoral condyles of rabbits [58]. They confirmed LacZ-positive cells in the transplanted area, but not elsewhere in the cartilage, for up to 4 weeks. Baragi et al. transplanted genetically engineered chondrocytes with adenovirus carrying the LacZ gene into full-thickness cartilage defects in combination with a type I collagen sponge as a scaffold [59]. They showed transient expression of -galactosidase limited to only 10 days. Recently, Ikeda et al. used the ex vivo approach to transplant adenovirally transduced chondrocytes embedded in type I collagen gels to osteochondral defects of rats [60]. They reported LacZ expression for up to 8 weeks, but the expression gradually decreased. Although the optimal period for transgene expression is still under investigation, further studies are necessary to achieve longer transgene expression. For transplanted cells, various types of cells can be used to engineer tissue for cartilage repair, from undifferentiated pluripotent cells, such as embryonic stem cells, to welldifferentiated cells, such as chondrocytes. Until now, chondrocytes have been extensively studied and are the natural, logical cell source for cartilage repair. However, for the purposes of a cell source and a vehicle to deliver therapeutic genes to the joints for cartilage defects, chondrocytes are not the only cells that can be used. Another candidate for a cell source and gene delivery vehicle is MDCs. In 1997, Day et al. showed feasibility of muscle cell–mediated ex vivo gene delivery to numerous intra-articular structures [61]. They injected primary muscle–derived cells and synoviocytes transduced by adenovirus carrying the gene encoding -galactosidase. They demonstrated that with gene delivery mediated by muscle-derived cells, LacZ was expressed in the synovial lining, meniscus, and cruciate ligament, whereas transduced synovial cells resulted in LacZ expression only in the synovium. Moreover, they showed that purified immortalized myoblasts fused more readily and resulted in more de novo intra-articular myofibers than primary myoblasts, indicating the importance of obtaining pure populations of myogenic cells, void of fibroblast and adipocyte contamination. With this ex vivo gene delivery, there is a great possibility that cartilage repair can be enhanced by transducing therapeutic growth factors. As for the cartilage defect model, in 2000, Lee et al. made full-thickness cartilage defects on rabbit femurs and treated them by transplantation in combination with muscle biopsy as a scaffold and autologous myoblasts as a cell source [62]. They showed favorable hyaline-like carti-

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lage formation, especially in the group that was supplemented by the administration of IGF-1 protein. VI ANTERIOR CRUCIATE LIGAMENT The anterior cruciate ligament (ACL) is a main restraint to anterior translation of the tibia and contributes mechanical stabilization of the knee. The ACL is the second most frequently injured ligament in the knee joint, following the medial collateral ligament (MCL). Considering the fact that MCL injuries do heal well spontaneously, the ACL has a very low healing capacity, because the ACL is an intra-articular ligament covered with a synovial sheath, and it is bathed in the synovial fluid. If ACL injuries are left untreated, they cause remarkable anterior and rotary instability of the joint that hinders the patient from participating in sports. Therefore, ACL reconstruction using autologous or allogenic tendons is usually recommended for patients with ACL injuries, especially active, young patients. Both the bone–patella tendon–bone (BPTB) and hamstring tendons, such as semitendinosus or gracilis tendons, are gold standards for ACL reconstruction. In the last two decades, the outcomes of ACL reconstructions have been considerably improved through biological and biomechanical research in this field. The ACL graft has been shown to undergo a ligamentization process comprising four phases: ischemic necrosis, revascularization, cell proliferation, and collagen remodeling [63]. However, there is still the challenge to improve and accelerate ligament healing after ACL reconstruction because it takes several months for the grafted tendons to achieve enough biomechanical strength. Even professional athletes are advised not to return to their previous sports activities until at least six months after ACL reconstruction. The ligaments consist of parallel collagen fibers, spindle cells, and extracellular matrix, such as dermatan–sulfate. Recently, some studies reported the beneficial effects of several growth factors on ACL fibroblasts. It was reported that epidermal growth factor (EGF), platelet-derived growth factor (PDGF), and bFGF can enhance the cell proliferation of ACL fibroblasts and that TGF- can increase proteoglycan synthesis and collagen synthesis [64–69]. Moreover, PDGF, TGF-, and EGF promote MCL healing in vivo [70,71]. These studies suggest that some growth factors may improve the healing of the ACL or ACL graft ligamentization. There are also the possibilities that the revascularization of the grafted tendons may be enhanced by VEGF and that reinnervation to the graft must be promoted by NGF. For the last two decades, many researchers have been focused on the biomechanical function of the ACL. The challenge has been to reconstruct a mechanically strong ligament that has normal biomechanical properties. However, many orthopedic surgeons have recognized the proprioceptive function of the ACL as well as the joint stabilizing function. Because the nerve supply of the ACL may contribute to the proprioceptive function, there is a possibility that NGF can enhance the proprioceptive function of the knee joint with a promotion of nerve growth to the grafted tendons after ACL reconstruction. In 1996, Nakamura et al. investigated the in vivo introduction of the LacZ gene into a healing rat patellar ligament using the hemagglutinating virus of Japan (HVJ) liposome–mediated gene transfer method [72]. They injected HVJ–liposome suspension containing -galactosidase cDNA into the injured site of the ligament 3 days after surgery. LacZ-positive cells were present for up to 56 days with a peak of expression at 7 days after gene transfer, indicating the feasibility of introducing a transgene into the cells of a healing ligament. In 1998, Nakamura et al. also investigated the effectiveness of HVJ–liposome suspension containing PDGF-B cDNA for ligament healing [73]. The HVJ-liposome sus-

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pension containing PDGF-B cDNA was injected directly into the injured patellar ligament of 14-week-old male Wistar rats. They reported that PDGF-B gene transfer caused the enhanced expression of PDGF in the healing ligament for up to 4 weeks after transfection, leading to an initial promotion of angiogenesis and subsequent enhanced collagen deposition in the wound. There are some reports of MDC-mediated gene transfer to the joint. As previously described, Day et al. showed the feasibility of MDC-mediated ex vivo gene delivery to numerous intra-articular structures. They injected primary MDCs and synoviocytes transduced by adenovirus carrying the LacZ gene, and demonstrated that with gene delivery–mediated MDCs, LacZ was expressed in the synovial lining, meniscus, and cruciate ligament [61]. In 1999, Menetrey et al. investigated the feasibility of myoblast-mediated gene transfer to the ACL [74]. Rabbit myoblasts and ACL fibroblasts were transduced with adenovirus carrying the LacZ reporter gene, and the genetically transduced myoblasts and fibroblasts were injected into the ACLs of adult rabbits. They reported that the injection of myoblasts into the ligament resulted in a higher number of cells expressing the marker gene than after the fibroblast injection. Beyond the first week, the formation of myotubes/myofibers was observed in the ligament as evidenced by desmin staining. They hypothesized that myoblasts would be capable of surviving under the unfavorable conditions of the ligament because of their capacity to fuse and become postmitotic. Another important aspect for ACL reconstruction is the tendon–bone healing in the femoral and tibial bone tunnels when we use hamstring tendons for the graft. Insufficient tendon–bone healing in the tunnels after ACL reconstruction causes instability of the joint as well as poor ligamentization of the graft. Bone morphogenetic protein-2 can enhance the healing between the grafted tendons and the surface of the bone tunnel. In 2001, Martinek et al. reported the technique of biological preconditioning of the ACL graft before ACL reconstruction [75]. They tested the feasibility of ex vivo gene transfer to the semitendinosus tendon graft after preconditioning with adenovirus in vitro. The semitendinosus tendons were harvested from mature New Zealand white rabbits and transduced with adenoviral vectors encoding the LacZ reporter gene. In vitro, the expression of -galactosidase to the superficial layer of the tendons was confirmed and declined gradually over 6 weeks. After performing the ACL reconstruction with genetically preconditioned semitendinosus tendons, two different patterns of transduction were observed [75]. In the intra-articular portions of the grafts, LacZ-positive cells were confirmed mostly along the surfaces of the tendons, and only a few scattered positive cells were observed in the deep layers of the tendons [75]. In the intratunnel portions of the ACL grafts, the number of the LacZ-positive cells did not decline during the observation time of 6 weeks. Moreover, a high number of transduced cells were found in the deep layers of the tendons, suggesting a migration of transduced cells from the tendon surfaces into the deep layers of the tendons. This study demonstrated that ex vivo transduction of a semitendinosus tendon before implantation to the joint can be used for biological preconditioning of the tendons and also the insertion site, showing the possibility of shortening healing time and strengthening the grafted tendon during its early remodeling phase [75]. VII MENISCUS The meniscus has important functions, such as load-bearing, shock absorption, load transmission, and joint stabilization. Meniscal injuries are very common in both younger and older generations. Since many reports revealed the high incidence of degenerative changes

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after total meniscectomy in younger patients, orthopedic surgeons have focused on preserving the meniscus as much as possible with meniscal repair techniques [76,77]. However, because only the outer one-third of the meniscus’s peripheral area is vascularized, its healing potential is very limited, especially in the inner two-thirds [78]. Of course, there is a possibility that meniscal healing can be enhanced by growth factors in combination with gene therapy. In 2000, Kasemkijwattana et al. characterized the effects of various doses of nine growth factors on the meniscal fibrochondrocyte proliferation and collagen and noncollagen synthesis [79]. It was reported that EGF, TGF-, bFGF, and PDGF-AB were candidate growth factors to improve meniscal healing. They also evaluated and showed the feasibility of direct and MDC-mediated ex vivo gene transfer into the meniscus. The persistence of -galactosidase expression was observed for up to six weeks for both direct and MDC-mediated ex vivo gene transfer to the injured meniscus. In 1999, Goto et al. investigated the feasibility of gene transfer in vitro and in vivo with a LacZ gene to the meniscus [80]. They demonstrated that gene expression persisted for at least 20 weeks under in vitro conditions. With regard to the in vivo study, gene expression persisted within the transplanted, transduced meniscal cells for at least 6 weeks, showing good feasibility of transferring exogenous genes to the meniscal injury site. In 2000, Goto et al. also transfected the human and canine meniscal cells with retroviruses carrying the gene encoding for the human TGF- cDNA and examined matrix synthesis in vitro [81]. This report revealed that meniscal cells are readily transduced by retroviral vectors and respond to the transfer of a TGF- cDNA by greatly increasing synthesis of matrix such as collagen or proteoglycan. VIII FUTURE DIRECTIONS BASED ON THE USE OF MULTIPOTENT STEM CELLS ISOLATED FROM SKELETAL MUSCLE We have also focused our research on stem cells isolated from muscle tissue. These cells are located in adult muscle tissue and are normally quiescent, but are activated in response to stress induced by weight bearing, exercise, or trauma. Applying muscle-derived progenitor cells combined with recent advances in gene therapy facilitates novel approaches to the treatment of musculoskeletal disorders. The approach of delivering growth factors ex vivo may revolutionize a musculoskeletal medical field that is historically based on biomechanical approaches. The advantage of muscle-based tissue engineering is the unique biology of skeletal muscle–derived cells. Skeletal muscle contains so-called satellite cells, which are resting mononucleated precursor cells. Data suggest that cells residing within skeletal muscle, like bone marrow–derived mesenchymal stem cells, can differentiate into several lineages [42]. Consequently, muscle-derived cells can regenerate various tissues. Cell isolation from skeletal muscle has already been well described [82,83]. With the help of the preplating technique for myoblast isolation, which has been previously described by Rando and Blau [84] and Qu et al. [85], we isolated different populations of muscle-derived cells that contained different ratios of desmin (myogenic specific marker) -positive and -negative cells [85]. The percentage of muscle-derived stem cells increased with the number of the preplate. Multipotent muscle-derived cells express myogenic markers and stem cell markers, such as Bcl-2, CD34, FLK-1, and Sca-1, but differ from mesenchymal and hematopoietic stem cells [42]. We found that 95% of the cells obtained and multiplied in a monolayer culture are desmin-positive and coexpress CD34 and Bcl-2 markers. In muscle tissue sections, the muscle origin of these cells has been established. Like satellite cells, Bcl-2- and CD34-positive cells are located beneath the basal membrane of myofibers [42].

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Clonal isolation and characterization from this CD34/Bcl-2 enriched population yielded a putative stem cell able to differentiate into other lineages (muscle, bone). Therefore, one future direction for the musculoskeletal system is muscle-derived stem cell–based ex vivo gene therapy in combination with tissue engineering. REFERENCES 1. Garrett W. E. 1996. Muscle strain injuries. Am. J. Sports Med. 24:S2–S8. 2. Jarvinen M., Lehto M. U. K. 1993. The effect of early mobilization and immobilization on the healing process following muscle injuries. Sports Med. 15:78–89. 3. Kasemkijwattana C., Menetrey J., Day C. S., Bosch P., Buranapanitkit B., Moreland M.S., Fu F. H., Watkins S. C., Huard J. 1998. Biological interventions in muscle healing and regeneration. Sports Med. Arthrosc. Rev. 6:95–102. 4. Menetrey J., Kasemkijwattana C., Fu F. H., Moreland M. S., Huard J. 1999. Suturing versus immobilization of a muscle laceration: a morphological and functional study in a mouse model. Am. J. Sports Med. 27:222–229. 5. Kasemkijwattana C., Menetrey J., Somogyi G., Moreland M. S., Fu F. H., Buranapanitkit B., Watkins S. C., Huard J. 1998. Development of approaches to improve the healing following muscle contusion. Cell Transplantation 7:585–598. 6. Kasemkijwattana C., Menetrey J., Bosch P., Somogyi G., Moreland M. S., Fu F. H., Buranapanitkit B., Watkins S. C., Huard J. 2000. Use of growth factors to improve muscle healing after strain injury. Clin. Orthop. Rel. Res. 370:272–285. 7. Menetrey J., Kasemkijwattana C., Day C. S., Bosch P., Vogt M., Fu F. H., Moreland M. S., Huard J. 2000. Growth factors improve muscle healing in vivo. J. Bone Joint Surg. Br. 82B:131–137. 8. Carlson B. M., Faulkner J. A. 1983. The regeneration of skeletal muscle fibers following injury: a review. Med. Sci. Sports Exercise 15:187–196. 9. Hughes C., Hasselman C. T., Best T. M., Martinez S., Garrett W. E. 1995. Incomplete, intrasubstance strain injuries of the rectus femoris muscle. Am. J. Sports Med. 23:500–506. 10. Nikolaou P. K., MacDonald B. L., Glisson R. R., Seaber A. V. Garrett W. E. 1987. Biomechanical and histological evaluation of muscle after controlled strain injury. Am. J. Sports Med. 15:9–14. 11. Hurme T., Kalimo H. 1992. Activation of myogenic precursor cells after muscle injury. Med. Sci. Sports Exercise 24:197–205. 12. Bischoff R. 1994. The satellite cell and muscle regeneration. In: Myology, 2nd ed. McGrawHill: New York, pp. 97–118. 13. Hurme T., Kalima H., Lehto H., Jarvinen M. 1991. Healing of skeletal muscle injury: an ultrastructural and immunohistochemical study. Med. Sci. Sports Exercise 23:801–810. 14. Lee C. W., Fukushima K., Usas A., Xin L., Pelinkovic D., Martinek V., Somogyi G., Robbins P. D., Fu F. H., Huard J. 2000. Biological intervention based on cell and gene therapy to improve muscle healing after laceration. J. Musculoskeletal Res. (4) 4:265–277. 15. Allamedine H. S., Dehaupas M., Fardeau M. 1989. Regeneration of skeletal muscle fiber from autologous satellite cells multiplied in vitro. Muscle Nerve 12:544–555. 16. Partridge T. A. 1991. Myoblast transfer: a possible therapy for inherited myopathies. Muscle Nerve 14:197–212. 17. Huard J., Labrecque C., Dansereau G., Robitaille L., Tremblay J. P. 1991. Dystrophin expression in myotubes formed by the fusion of normal and dystrophic myoblasts. Muscle Nerve 14:178–182. 18. Karpati G., Pouliot Y., Zubrzycka-Gaarn E. E., Carpenter S., Ray P. N., Worton R. G., Holland P. 1989. Dystrophin is expressed in mdx skeletal muscle fibers after normal myoblast implantation. Am. J. Pathol. 135:27–32.

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25 New Methods to Enhance the Regeneration of Muscle Jacques Ménétrey and Charles R. Bader University Hospital of Geneva, Geneva, Switzerland Laurent Bernheim University of Geneva, Geneva, Switzerland Johnny Huard University of Pittsburgh, Pittsburgh, Pennsylvania

I INTRODUCTION Muscle injuries are common; their incidence varies from 10 to 55% of all injuries sustained in sports [1]. In minor or major trauma, muscle tissue is always involved and muscle injuries are regularly associated with fractures. Muscle injuries are divided into two types: a shearing injury, in which both the myofibers and the connective tissue framework are torn, or an in situ injury, in which only the myofibers are damaged. In the latter type of injury, the basal lamina and connective tissue sheaths do not undergo significant damage. Shearing injuries, the most frequent muscle injury in sports, may be lacerations, contusions, or strains depending upon the mechanism of injury [1]. Contusion occurs when a muscle is injured by a significant compressive force, such as a direct blow. Such a mechanism is common in contact sports. A strain occurs when a forceful eccentric contraction is applied to an overstretched muscle, especially in jumping or sprinting [1,2]. This injury is common near the musculotendinous junction (MTJ) of a superficial muscle which crosses two joints, such as the rectus femoris, semitendinosus, and gastrocnemius muscles. Though rather rare in sports, muscle laceration is a dramatic injury, which consistently incapacitates athletes for long periods of time and often jeopardizes their professional careers. 499

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During their practice, orthopedic surgeons have to face two other challenging muscle conditions: compartment syndrome and limb lengthening. Compartment syndrome is characterized by an increased pressure within a closed compartment bounded by bone and fascia and leads to muscle damage. The increased pressure diminishes tissue perfusion, which may lead to an ischemic injury [3]. Common causes of increased compartment pressure include bleeding into a compartment, edema following partial or temporary ischemia, and crushing injuries. Ischemic contracture and nerve damage constitute the major consequences of inadequately treated compartment syndromes. During limb lengthening, the muscle is unable to keep pace with the elongation and this results in contracture, despite its high regenerative capabilities. II MUSCLE HEALING PROCESS Muscle regeneration and healing depend heavily upon skeletal muscle satellite cells. Satellite cells were first described in frog muscle by Mauro [4] and have since been observed in adult avian and mammalian muscle (reviewed by Bischoff [5]). Satellite cells are small mononucleated fusiform cells lying between the basal lamina and the sarcolemma of muscle fibers. These cells are normally mitotically quiescent, but they can be activated during postnatal muscle growth [6] or when there is muscle damage. Activated satellite cells give rise to myoblasts that proliferate for a few days and then fuse to form new myotubes [5]. Myoblasts can also fuse with existing myofibers. It seems valid to consider satellite cells as myogenic stem cells. Indeed, they give rise to a population of proliferating myoblasts, but at the same time they renew their own reservoir. This renewal ability is confirmed by the observation that the number of quiescent satellite cells in adult muscle remains relatively constant over several cycles of degeneration and regeneration [7–9]. This capacity, however, is not infinite. In particular it is markedly reduced in severe muscle diseases such as Duchenne muscular dystrophy [10]. There has been considerable progress in understanding the molecular mechanisms that regulate the activation of muscle satellite cells, but a full review of the factors involved in this process is beyond the scope of this chapter. A recent review [11] addresses the issue of the activation and differentiation of satellite cells, as well as their embryonic origin. We shall simply note that recent work has identified new markers of satellite cells, in particular CD34 (a marker of stem cells and early progenitors of the hematopoietic system) and Myf5 (an early marker of myogenic commitment) [12]. Quiescent satellite cells also express the c-Met receptor tyrosine kinase, which is the receptor for hepatocyte growth factor, a factor released following a lesion and which triggers satellite cell activation. There are interesting and somewhat controversial data regarding the role of MyoD and Myf5 in the control of the self-renewal of satellite cells and in the amplification of their progeny, i.e., the myoblasts that proliferate to repair muscle damage [11,12]. Finally, it is important to mention that during the regeneration process satellite cells and myoblasts are influenced by environmental factors released by damaged muscle and by inflammatory cells, such as polymorphonuclear lymphocytes and macrophages, as well as by the proliferating myoblasts themselves (via autocrine mechanisms). These environmental factors include fibroblast growth factor, insulin-like growth factors, leukemia inhibitory factor, and interleukin-6 (see review by Seale and Rudnicki [11]). III PURIFICATION OF HUMAN MUSCLE SATELLITE CELLS In view of the role of satellite cells in regeneration and repair, techniques for purifying these cells and modifying their properties may have important clinical applications. We shall de-

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scribe methods we recently developed to purify satellite cells and self-renewing myogenic stem cells born in culture and examine ways of transfecting these purified cells. A Purification of Human Satellite Cells Methods for the purification of human satellite cells have been described by Blau and Webster [13] and by Webster et al. [14]. We have developed a simple procedure for purifying satellite cells directly from freshly dissociated human skeletal muscles [15] which is based on the cellular properties of satellite cells and does not require labeling cell surface markers. Microscopic observations of cellular suspensions obtained from freshly dissociated muscles reveal the existence of two types of cells, small cells (8–12 m in diameter) and large cells (15–20 m in diameter). By manually separating the two types of cells and by clonally culturing them we have been able to demonstrate that only small cells corresponded to myogenic human muscle satellite cells. They proliferate as myoblasts in a culture medium favoring proliferation and give rise to myotubes when cultured in a differentiation medium [15]. This observation was confirmed by flow cytometry analysis, which, in addition, revealed differences in the nucleocytoplasmic ratio of the small and large cells. We were able to perform flow cytometry sorting with biopsies weighing as little as 300 mg, which corresponded to a total cell number of about 30,000 in the dissociated cell suspension. The yield of satellite cells by flow cytometry was approximately 40 cells per milligram muscle. We have recently adapted these technique to small biopsy samples (10 mg) obtained with needle biopsies [16]. B Identification and Purification of Self-Renewing Myoblasts in the Progeny of Single Human Muscle Satellite Cells We have demonstrated that self-renewing myoblasts can be identified in the progeny of single human muscle satellite cells in culture [17]. Using cytoskeletal proteins and cell size as markers, we were able to demonstrate that self-renewing myoblasts were phenotypically different from other proliferating myoblasts, but similar to native satellite cells. We found that these presumptive myogenic stem cells born in culture are already segregated during myoblast proliferation. In contrast to the vast majority of the myoblasts that proliferate in culture, they are small and do not divide, as noted by the observation that they do not take up [3H]-thymidine. However, when they are isolated from their sister cells and cultured as single cells, they become activated and give rise to a progeny of myoblasts which are able to fuse. This contrasts with the population of larger cells found among proliferating myoblasts, which take up [3H]-thymidine but are unable to produce clonal cultures [17]. Enriched preparations of cells with myogenic stem cell–like properties can be obtained from proliferating myoblast cultures by flow cytometry on the basis of size and nucleocytoplasmic ratio. As proposed recently by Beauchamp et al. [18], this population of myogenic stem cells may correspond to the 2% of transplanted cells that survive the transplantation procedure. IV CLINICAL TREATMENT The current treatment of muscle injuries has essentially remained unchanged for decades. The immediate care consists of protection, rest, ice, compression, and elevation (PRICE). The aim is to prevent hematoma formation and interstitial edema, thus decreasing tissue ischemia. Nonsteroidal anti-inflammatory medication should be instituted only in the early

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phase because their long-term use may be detrimental [19]. Glucocorticoids should not be used since they delay the elimination of hematoma and necrotic tissue and delay muscle regeneration [20]. After 3 or 4 days, physical therapy consisting of stretching, strengthening, and ultrasound should be instituted. However, the use of ultrasound has yet to be proven in this setting. The final rehabilitation phase is sport-specific training. V NEW METHODS TO ENHANCE THE HEALING OF MUSCLE INJURY A Autologous Myoblast Transplantation Autologous myoblast transplantation (AMT), the implantation of myoblast precursors (satellite cells), has been extensively studied as a method to promote muscle regeneration and create a reservoir of normal myoblasts enabled to fuse and deliver genes to skeletal muscle. The potential of myoblast transplantation has also been investigated in the management of Duchenne muscular dystrophy (DMD) [21]. This approach can deliver and restore structural protein, such as dystrophin in DMD muscle and permits an increase in strength. Autologous myoblast transplantation can be utilized in an attempt to improve muscle healing following an injury. A muscle biopsy from a noninjured muscle of the same individual serves as an autologous donor. The myoblasts are isolated by enzymatic digestion from the biopsy material, cultured in an enriched milieu, and then injected into the damaged muscle. The major hurdle of myoblast transplantation is the well-documented immune rejection phenomenon [22], which may be circumvented by using autologous myoblast transfer. Studies on DMD have revealed that this approach allows the delivery of genes, the improvement of muscle regeneration, and the enhancement of strength in dystrophic muscle. Therefore, autologous myoblast transplantation may promote muscle regeneration and improve muscle healing following a severe injury. In a previous study, we have demonstrated enhancement of muscle regeneration after autologous myoblast transplantation in a muscle that had been injured by myonecrotic agents [23]. We observed that more than 90% of the transplanted muscle was populated with myofibers formed by the fusion of the injected myoblasts [23]. These participated to muscle regeneration by fusing with existing necrotized myofibers and thus prevented the muscle fibers from complete degeneration [23]. Furthermore, the ability of these myoblasts to secrete trophic substances which are critical for the regeneration process might stimulate the entire repair process and improve muscle healing. Most of the recent research has been focused on using myogenic precursor cell. These cells are isolated following a muscle biopsy (the different techniques are described in another section) amplified in vitro and injected into a muscle lesion. However, their effect on muscle healing is still under investigation. B Growth Factors and Muscle in Vitro Recently, new substances, such as growth factors, have been found to play a determinant role in muscle regeneration and healing [24,25]. Growth factors are small peptides that bind to membrane receptors to influence the various steps of the growth and development of cells via several signaling pathways [24,26]. It has already been shown that growth factors are capable of stimulating the growth and protein secretion of many musculoskeletal cells [27]. During muscle regeneration, it is presumed that trophic substances released by the injured muscle activate the satellite cells [5,8,28–30]. During growth and development, many growth factors have been shown to be capable of eliciting varying responses from skeletal

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muscle [24,26,31]. There has been some preliminary data suggesting that individual growth factors play a specific role during muscle regeneration [24–26,32–34]. In vitro studies have demonstrated that basic fibroblast growth factor (bFGF) stimulates cell proliferation while inhibiting differentiation in bovine and chick myoblasts in culture [35,36]. The mechanism for bFGF-stimulated proliferation appears to be by advancing cells from G0 to G1 in the cell cycle [24,31]. More recently, bFGF has been found to be capable of stimulating proliferation and repressing differentiation in MM14 mouse myoblast [37]. This inhibition of differentiation is thought to be due to repression of the “cell commitment” (the irreversible fate of a particular cell lineage to differentiate). However, in vitro, insulin growth factor type 1 (IGF-1) has been found capable of powerfully stimulating myoblast proliferation and differentiation [31,38–40]. In the Growth and Development Laboratory at the University of Pittsburgh, we have also investigated the effect of various growth factors on myoblast proliferation and differentiation in vitro. Myoblasts have been cultured at different concentrations (1, 10, and 100 ng/mL) with basic and acidic fibroblast growth factor (bFGF, aFGF), insulin growth factor type 1, nerve growth factor (NGF), platelet-derived growth factor-AA (PDGF-AA), transforming growth factor- and - (TGF-, TGF-). The myoblast proliferation and differentiation have been monitored at 48 and 96 h postincubation. We have observed that IGF-1, bFGF, and NGF were potent stimulators of both myoblast proliferation and differentiation in vitro (Table 1) [41]. Interestingly, this stimulation was dose-dependent [41]. The other growth factors showed no stimulating effect on myoblast proliferation and differentiation. These results suggest that bFGF, IGF-1, and NGF enhance both myoblast proliferation and differentiation and are the logical choice for being delivered to the site of a muscle injury to improve healing. C Growth Factors and Muscle in Vivo Despite the experimental elegance of the in vitro system it is important to recognize that regeneration in vivo is more complex due to the involvement of circulatory and intercellular communication [24,26]. There has been, however, some preliminary characterization of Table 1 Growth Factor Effect on Myoblast Proliferation and Fusion in vitro Growth factor

Proliferation

Fusion

bFGF IGF-1 NGF aFGF PDGF-AA EGF TGF- TGF-

Stimulatesa Stimulatesa Stimulatesa Inhibits Inhibits Inhibits Inhibits Inhibits

Stimulatesa Stimulatesa Stimulatesa Stimulatesa Inhibits Inhibits Inhibits Inhibits

ANOVA, p  0.05. Note: bFGF, basic fibroblast growth factor; IGF-1, insulin like growth factor type 1; NGF, nerve growth factor; aFGF, acidic fibroblast growth factor; PDGF-AA, platelet-derived growth factor AA form; EGF, epidermal growth factor; TGF-, transforming growth factor-a; TGF-, transforming growth factor-. a

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the role of certain growth factors during muscle regeneration which suggests that the individual roles of growth factors are similar to their individual effects seen in vitro [25,42]. Other studies have shown that during the muscle regeneration process following any injury, bFGF is present in the extracellular space as early as 8 h after the injury. It reaches a peak at 24 h, with the levels slowly decreasing over a period of 1 week [43]. Insulin growth factor type 1 is present after 2 days, reaches its peak at 3 days, and decreases over a period of 1 week [33,34]. Based on our in vitro study, we have injected bFGF, IGF-1, and NGF into a mouse muscle laceration and monitored the muscle regeneration by regular and quantitative histology 1 week postinjury. At 1 month postinjury, injured muscles were histologically assessed, and contractile properties were determined. This study has shown that bFGF, IGF1, and to a lesser extent NGF, improved muscle regeneration in mouse muscle (Fig. 1). We

Figure 1 Hematoxylin-eosin staining of the laceration site at 7 days postinjury. (A, C, E) Control with regenerating myofibers in the deep part of the muscle and infiltration of inflammatory cells in the superficial part. (B) Laceration injected with NGF, (D) bFGF and (F) IGF-1. Regenerating myofibers are located throughout the injured site (deep and superficial area). (Magnification 10.)

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Figure 2 Hematoxylin-eosin staining of the laceration site 1 month after the injury. (A, C, E) Control with a superficial layer of fibrotic tissue and small regenerating myofibers and a deep layer containing and numerous regenerating myofibers. (B) Laceration injected with NGF; (D) bFGF and (F) IGF-1. The appearance of the muscle injected with NGF is similar to control. The muscles treated with bFGF and IGF-1 show large regenerating myofibers filling the entire laceration site. (Magnification 10.) have documented an increase in the number and size of the regenerating myofibers, which is an index of muscle regeneration. In addition, our laceration muscle model has shown that regenerating myofibers were located in the superficial area of the injured site only when treated with growth factors, thus demonstrating greater initial muscle healing when the injured muscle is treated with specific growth factors [44]. At 1 month, the nontreated muscle showed numerous centronucleated regenerating myofibers in the deepest part of the laceration (Fig. 2). Superficially, the laceration was covered with fibroblastic tissue in which many regenerating myofibers were found. In the muscle treated with IGF-1 and bFGF, the regenerating myofibers were uniformly located

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in the deep and superficial part of the muscle. Their diameter was similar to the surrounding normal myofibers, and many of their nuclei were already peripherally located. The development of fibroblastic tissue was also reduced in the treated muscle. These findings suggest that the muscle healing was accelerated in these muscles when compared to the control. The muscle treated with NGF contained numerous mononucleated regenerating myofibers in the deep part of the muscle. In the superficial part of the muscle, regenerating myofibers of small diameter were observed as well as areas of fibroblastic tissue. These muscles had the same histological appearance as the control (Fig. 2). One month after the injury, the twitch and tetanus strengths were increased in muscles treated with IGF-1 and bFGF. To minimize interanimal variation, the data were normalized with respect to untreated controls, i.e., the strength in the experimental muscle was divided by that in the control muscle in the contralateral side and multiplied by 100 to determine the percentage of change. The fast twitch strength was increased by 76 (14%) (p0.001) for bFGF and 164 (36%) (p0.005) for IGF-1 when compared to control. The tetanus strength was increased by 74 (20%) (p0.002) and 106 (34%) (p0.003) [44]. The muscles treated with NGF regularly produced a mean twitch and tetanus strength inferior to the controls, but the difference was not statistically significant. Therefore, it appears that the administration of exogenous growth factor is a promising approach to improve muscle healing. However, the successful clinical implementation of this technique is currently limited by the problem of maintaining an adequate concentration of growth factor in the lesion site or target tissue. The short half-life of growth factors and systemic lavage may lead to a rapid clearance of the substances from the desired site. To address these issues, gene therapy may be an interesting delivery system to the muscle. VI GENE THERAPY IN MUSCLE New delivery techniques are required to achieve sustained and efficient local delivery of therapeutic proteins such as growth factors. The new delivery system should aim at circumventing the rapid clearance and the nonspecificity of the growth factor action, as well as allow the release of an efficient concentration of protein. Gene therapy provides a promising approach to meet these requirements. The genetic information (usually a cDNA) encoded for the therapeutic protein is inserted into living cells, using nonviral and viral vectors. The genetically modified cells express the protein encoded by the transferred DNA in a sustained manner. With such a technology the growth factor can be delivered to the tissue locally on a mid- to long-term basis, thus avoiding repeated injections and systemic administration. Efficient gene transfer is the first step to address in order to achieve a successful delivery system. Extensive work from our laboratory and others has been directed toward promotion of muscle growth and enhancement of muscle regeneration, specifically aimed at reducing the muscle weakness, secondary to such muscle diseases as Duchenne muscular dystrophy. The different gene transfer systems developed for these muscle diseases may also be useful for other muscle disorders. Recent effort has been directed at the development of vectors for efficient gene delivery to muscle. Plasmid DNA, liposomes, and viral vectors have been used to transfer genes into skeletal muscle. Direct transfection of naked DNA into muscle cells in vitro as well as in vivo initially had been found to be inefficient [45]. However, transgene expression has been found to persist for up to 1 year, which suggests low immunogenicity and cytotoxicity related to the use of nonviral vectors [46,47].

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A Viral Vector for Muscle Retrovirus, adenovirus, and HSV-1 have also been investigated as viral gene delivery vehicles to skeletal muscle. These viral vectors were found capable of highly transducing muscle cells, but were hindered by several limitations including a differential viral transduction throughout muscle maturation, transient transgene expression related to cytotoxicity, and immunological problems against the viral vectors [47–50]. Recently, adeno-associated virus (AAV) has been investigated as another viral gene delivery vehicle to skeletal muscle. Since this viral vector bypasses some of the limitations associated with the currently used gene transfer systems to skeletal muscle, it may become an efficient vehicle to deliver therapeutic protein to the injured muscle [51–53]. Direct gene therapy, which consists of directly injecting the vectors into muscles, has been extensively used to mediate gene delivery to skeletal muscle. Direct gene therapy approaches based on naked DNA [45], retrovirus [49], adenovirus [48,54,55], herpes simplex virus type 1 [56,57], and adeno-associated virus [53,58] have been designed to deliver genes to skeletal muscle. With vectors carrying reporter genes, muscle cells have been successfully transduced in vitro as well as in vivo using replication-defective adenovirus, retrovirus, herpes simplex virus, and adeno-associated virus recombinants. Although a poor level of gene transfer has been observed with many of these viral vectors in mature skeletal muscle, some of them, such as adenovirus and adeno-associated virus, allow an efficient gene transfer in adult regenerating muscle. Adenovirus and adeno-associated virus are capable of infecting dividing and nondividing cells and appear to be the vectors of choice utilizing the direct approach. Another way of gene delivery is the ex vivo approach. This consists of establishing a primary myoblast cell culture from injured muscle and infecting them with engineered vectors. Then the transduced cells are injected into the same host. This method has already been performed with the use of recombinant adenovirus [59,60], retrovirus [61], and herpes simplex virus [62] carrying reporter genes (-galactosidase or luciferase). Several studies have shown that the transduced myoblasts (isogenic myoblasts) fused and introduced reporter genes into the injected muscle [59,61,62]. Furthermore, a higher efficiency of gene transfer has been found using the ex vivo approach versus the direct injection of the same amount of virus [60,62]. Since myoblast transplantation and gene therapy have been primarily hindered by immunorejection, the ex vivo gene transfer offers the advantage of reducing immunological problems against the injected myoblasts. This statement is not valid for the adenoviral vector, which expresses viral particles at the transduced cell surface and induces an important immunological response. For the ex vivo approach, retrovirus and adeno-associated virus appear to be the vectors of choice. For orthopedic applications, however, it is possible that nonviral vectors (plasmid DNA, liposomes) will permit a sufficient level of gene transfer to achieve an efficient delivery system. B Efficient Nonviral DNA-Mediated Gene Transfer to Human Primary Myoblasts Using Electroporation Recent studies on myoblast transplantation indicate that both inflammatory and immune processes contribute to the rapid death of transplanted myoblasts and that all myoblasts may not be equally efficient for transplantation [18,63]. This may explain the failures of the first clinical trials in Duchenne muscular dystrophy patients. Several strategies have been proposed to improve the survival of transplanted myoblasts. One is to provide immunosuppression and/or anti-inflammatory agents by gene therapy since myoblasts can be genetically engineered to secrete TGF- [64], inhibitors of interleukin-1 (IL-1) [63] or

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inhibitors of IL-12 [62]. A second technique is to use myoblasts with stem cell–like properties, which have recently been shown to better survive transplantation [18]. Engineering myoblasts with stem cell–like properties would thus be of particular interest in order to reach optimal results. Successful gene therapy with genetically engineered myoblasts implies efficient and reproducible gene transfer into myogenic cells. Viral-mediated systems are very efficient, but suffer from limited capacity for foreign DNA and clinical use raises important safety issues [66]. Approaches to enhance the efficiency of nonviral DNA-mediated gene transfer systems have included particle bombardment [67], DNA precipitation [68], and, more recently, cationic liposome vesicles [69,70]. Although these transfection procedures yield encouraging results with muscle cell lines, they are not very efficient in human primary myoblasts and do affect cell viability [69–72]. We have recently obtained encouraging results with transfection by electroporation of primary culture of human myoblasts [73]. Electroporation is a physical process that transiently permeabilizes cell membranes with an electrical pulse allowing the uptake of a wide variety of biological molecule including dyes or drugs. We have shown that by careful selection of the parameters, a relatively narrow window can be found which allows a high and reproducible gene transfer into cultured human primary myoblasts combined with good cell survival in culture. We have also shown that electroporated myoblasts are morphologically and functionally indistinguishable from untransfected myoblasts, express developmentally regulated ionic currents, and fuse to form myotubes. In addition, we found that the subpopulation of human myoblasts with self-renewing properties (see previous discussion) can be as efficiently transfected by electroporation as proliferating myoblasts [73]. VII POTENTIAL APPLICATION IN THE ORTHOPEDIC FIELD The tissues which form the musculoskeletal system exhibit different healing capacities following an injury and, with the exception of bone, there is no restitutio ad integrum [74]. The healing process in soft tissue always ends with the formation of scar tissue, which has lower mechanical and functional properties than the original tissue. Some tissues have low healing capacities, such as articular cartilage, the meniscus, and the anterior cruciate ligament. Others, such as muscle and bone, have high healing capacities. Despite a high capacity for regeneration, the response of muscle to serious injury typically involves the formation of dense fibrotic scar tissue interposed between normal muscle. Enhancing muscle growth and regeneration may possibly prevent the formation of dense scar tissue and thus improve the quality of muscle healing. This might reduce the risk of reinjury and decrease the incidence of muscle pain secondary to scarring. The enhancement of muscle growth and regeneration might also prevent the occurrence of contractures secondary to limb lengthening and/or compartment syndrome. Gene therapy is not yet an established therapeutic method in the treatment of orthopedic disorders or sports injuries. However, we believe that gene therapy has great potential as a delivery system to musculoskeletal tissue because it allows a continuous local delivery of a therapeutic protein, such as growth factor. The feasibility of gene transfer to musculoskeletal tissues has recently been investigated and showed promising results. Over the last several years our research center has studied the feasibility of gene transfer to muscle injury. After the development in mice of muscle strain, contusion, and laceration models, direct and ex vivo approaches of gene transfer have been performed on the different injured muscles. Either 10 L of 2.5  106 recombinant adenovirus carrying the LacZ reporter gene or 10 L containing 1  106 transduced myoblasts (MOI25) was injected

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into injured muscles. The muscles were harvested 1 week later, cryosectioned, and assayed for -galactosidase expression. With the use of both direct and ex vivo gene transfer, we have observed the presence of many -galactosidase–positive myofibers in the strained, contused, and lacerated muscles, which had been injected with the recombinant adenovirus or the transduced myoblasts (Fig. 3). These transduced myofibers were located at the injured site.

Figure 3 Adenovirus-mediated direct and ex vivo gene transfer in a (A, B) strained, (C, D) contused, and (E, F) lacerated muscle. (A, C, E) We observed many transduced myofibers at the injured sites following direct injection of adenoviral vector at 7 days postinjury (B, D, F). We detected also LacZ-positive myofibers at the injured sites following ex vivo gene transfer of the -galactosidase reporter gene at 5 days postinjection. (B, D, F) The presence of these transduced myofibers suggested that our injected myoblasts have fused into the injured area mediating gene transfer in injured muscle. (Magnification 10.)

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In summary, both direct gene transfer and ex vivo gene delivery mediated by adenovirus have been found capable of delivering the -galactosidase reporter gene to the injured muscle at 5 days post-trauma. The development of an approach based upon the use of gene therapy to deliver growth factors aimed at a quicker and more complete recovery may revolutionize the significant “downtime” following a muscle injury. Indeed, the enhancement of muscle growth and regeneration by continuous and local delivery of specific growth factors may limit the formation of scar tissue, accelerate muscle recovery, and result in a complete healing after severe muscle strain, laceration, and contusion. Furthermore, this may possibly reduce the formidable risk of reinjury at the junction between scar tissue and regenerated muscle, thus preventing recurrences of muscle injury which jeopardize an athlete’s carrier. Gene therapy must remain a therapeutic tool and not be used as a new means to enhance performance. The entire scientific community must be concerned by this issue and act responsibly at this early stage since the new technology will be widely available at the preclinical level. Scientists involved in gene therapy projects related to sports medicine should form an ethics committee with its primary task to regulate the application of these new techniques. There is a real danger that the misuse of gene therapy will lead to an unfair manipulation of sports performance and expose athletes to hazardous secondary health effects. The development of the use of gene therapy to deliver growth factors may be very helpful in certain muscle conditions, such as with the secondary effect of limb lengthening, compartment syndrome and ischemia, and muscular dystrophy. The regenerative capabilities of muscle is illustrated by the increased satellite cell release in elongated muscle [75]. However, the regenerative response is often inadequate as the muscle is unable to keep pace with the limb lengthening and results a contracture. Enhancing muscle growth and regeneration would allow the elongated muscle to follow the lengthening of the limb without structural damage, and the occurrence of a contracture would be prevented. Compartment syndrome and ischemia result in conditions that are different from limb lengthening. Following an ischemic trauma, the muscle is partially or completely necrotic and the regenerative capacity initially depends upon the numbers of living myogenic precursor cells. With regeneration and repair the gene transfer technique, with its potential for continuously expressing and delivering an exogenous source of growth factors to the injured muscle, may promote healing and limit the formation of connective tissue scar. The application of gene therapy technology to the muscular dystrophy diseases, especially the Duchenne type, represents one of the most challenging fields of investigation in orthopedic medicine. In this situation, one needs to deliver the therapeutic protein (dystrophin) to all the muscles of the organism, particularly the respiratory muscles such as the diaphragm. These are not easily accessible and represent an important muscle mass to be transduced. Furthermore, the gene expression must persist for the patient’s entire life to be definitely therapeutic, in comparison to a short-term persistence which would be sufficient to create a significant improvement in a sports medicine application. In conclusion, we believe that gene therapy has great potential as a delivery system to muscle tissue. Both in vitro and in vivo studies have demonstrated that specific growth factors are potent stimulators of muscle cell proliferation and differentiation. The continuous and local production of these proteins by gene delivery in an injured muscle represents a promising and revolutionary way of treating severe injury. This method may have the potential to change the treatment and outcome of severe muscle strain, extensive laceration,

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surgical trauma, and large and multiple contusions. In addition, it promises to offer new insight into the treatment of other muscle disorders, such as fibrosis after limb lengthening, compartment syndrome, and muscular dystrophy. ACKNOWLEDGMENTS The authors wish to thank Richard Stern, M.D., Clinique d’Orthopedie, University Hospital of Geneva, Geneva, Switzerland, for his assistance in the preparation of the manuscript and Marcelle Pellerin and Ryan Prucnic, Growth and Development Laboratory, University of Pittsburgh Medical Center, for their technical assistance. REFERENCES 1. Lehto M., Jarvinen M. 1991. Muscle injuries healing and treatment. Annales Chirurgiciae et Gynaecologiae 80:102–109. 2. Garrett Jr. W. E. 1990. Muscle strain injuries: clinical and basic aspects. Med Sci Sports Exerc, 22:436–443. 3. Mubarak S., Hargens A. R. 1981. Compartment syndromes and Volksmann’s contracture. WB Saunders: Philadelphia, pp. 106–118. 4. Mauro A. 1961. Satellite cells of skeletal muscle fibers. J. Biophys. Biochem. Cytol. 9:493–495. 5. Bischof R. 1994. The satellite cell and muscle regeneration. In: Myology, Engel, A. G., Franzini-Armstrong C., ed. McGraw-Hill: New York, pp. 97–118. 6. Moss F. P., Leblond C. P. 1971. Satellite cells as the source of nuclei in muscles of growing rats. Anat. Rec. 170:421–435. 7. Gibson M. C., Schultz E. 1983. Age-related differences in absolute numbers of skeletal muscle satellite cells. Muscle Nerve 6:574–580. 8. Schultz E., Jaryszak D. L., Valliere C. R. 1985. Response of satellite cells to focal skeletal muscle injury. Muscle Nerve 8:217. 9. Schultz E., Jaryszak D. L. 1985. Effects of skeletal muscle regeneration on the proliferation potential of satellite cells. Mech. Ageing Dev. 30:63–72. 10. Webster C., Blau H. M. 1990. Accelerated age-related decline in replicative life-span of Duchenne muscular dystrophy myoblasts—implications for cell and gene therapy. Somatic Cell Mol. Genet. 16:557–565. 11. Seale P., Rudnicki M. A. 2000. A new look at the origin, function, and “stem-cell” status of muscle satellite cells. Devel. Biol. 218:115–124. 12. Beauchamp J. R., Heslop L., Yu D. S., Tajbakhsh S., Kelly R. G., Wernig A., Buckingham M. E., Partridge T. A., Zammit P. S. 2000. Expression of CD34 and Myf5 defines the majority of quiescent adult skeletal muscle satellite cells. J. Cell Biol. 151:1221–1234. 13. Blau H. M., Webster C. 1981. Isolation and characterization of human muscle cells. Proc. Natl. Acad. Sci. USA 78:5623–5627. 14. Webster C., Pavlath G. K., Parks D. R., Walsh F. S., Blau H. M. 1988. Isolation of human myoblasts with the fluorescence-activated cell sorter. Exper. Cell Res. 174:252–265. 15. Baroffio A., Aubry J. P., Kaelin A., Krause R. M., Hamann M., Bader C. R. 1993. Purification of human muscle satellite cells by flow cytometry. Muscle Nerve 16:498–505. 16. Magistris M. R., Kohler A., Pizzolato G., Morris M. A., Baroffio A., Bernheim L., Bader C. R. 1998. Needle muscle biopsy in the investigation of neuromuscular disorders. Muscle Nerve 21:194–200. 17. Baroffio A., Hamann M., Bernheim L., Bochaton-Piallat M. L., Gabbiani G., Bader C. R. 1996. Identification of self-renewing myoblasts in the progeny of single human muscle satellite cells. Differentiation 60:47–57.

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40. Florini J. R., Ewton D. Z., Roof S. L. 1991. Insulin-like growth factor-1 stimulates terminal myogenic differentiation by induction of myogenin gene expression. Mol. Endocrinol. 5:718. 41. Menetrey J., Kasemkijwattana C., Day C. S., et al. 1998. Characterization of trophic factors to promote muscle growth. Trans. Orthop. Res. Soc. 42. DiMario J., Buffinger N., Yamanda S., et al. 1989. Fibroblast growth factor in the extracellular matrix of dystrophic (mdx) mouse muscle. Science 244:688–690. 43. Anderson J. E., Mitchell C. M., McGeachie J. K., Grounds M. 1995. The time of basic fibroblast growth factor expression in crush-injured skeletal muscles of SJL/J and BALB/c mice. Exper. Cell Res. 216:325–334. 44. Menetrey J., Kasemkijwattana C., Day C., Bosch P., Moreland M., Fu H. F., Huard J. 2000. Growth factors improve muscle healing in vivo. J. Bone Joint Surg. 82B:131–137. 45. Ascadi G., Dickson G., Love D., et al. 1991. Human dystrophin expression in mdx mice after intramuscular injection of DNA constructs. Nature 352:815–818. 46. Katsumi A., Emi N., Abe A., et al. 1994. Humoral and cellular immunity to an encoded protein induced by direct DNA injection. Hum. Gene Ther. 5:1335–1339. 47. Wolfe J. A., Ludkte J. J., Acsadi G., et al. 1992. Long term persistence of plasmid DNA and foreign gene expression in mouse muscle. Hum. Molec. Genet. 1:363–369. 48. Ascadi G., Jani A., Huard J., et al. 1994. Cultured human myoblasts and myotubes show markedly different transducibility by replication-defective adenovirus recombinants. Gene Ther. 1:338–340. 49. Dunkley M. G., Wells D. J., Walsh F. S., et al. 1993. Direct retroviral-mediated transfer of a dystrophin minigene into mdx mouse muscle in vivo. Hum. Mol. Genet. 2:717–723. 50. Huard J., Akkaraju G., Watkins S. C., et al. 1997. Lac-Z gene transfer to skeletal muscle using a replication-defective herpes simplex virus type 1 mutant vector. Hum. Gene Ther. 8:439–452. 51. Clark K. R., Sferra T. J., Johnson P. R. 1997. Recombinant adeno-associated viral vectors mediate long term transgene expression in muscle. Hum. Gene Ther. 8:659–669. 52. Fisher K. J., Jooss K., Alston J., et al. 1997. Recombinant adeno-associated virus for muscle directed gene therapy. Nat. Med. 3:306–312. 53. Xiao X., Li J., Samulski R. J. 1996. Efficient long-term gene transfer into muscle tissue of immunocompetent mice by adeno-associated virus vector. J. Virol. 70:8098–8108. 54. Huard J., Lochmueller H., Acsadi G., et al. 1995. Differential short-term transduction efficiency of adult versus newborn mouse tissues by adenoviral recombinants. Exper. Molec. Pathol. 62:131–143. 55. Huard J., Lochmueller H., Acsadi G., et al. 1995. The route of administration is a major determinant of the transduction efficiency of rats tissue by adenoviral recombinants. Gene Ther. 2:107–115. 56. Huard J., Goins B., Glorioso J. C. 1995. Herpes simplex virus type 1 vector mediated gene transfer to muscle. Gene Ther. 2:1–9. 57. Huard J., Akkaraju G., Watkins S. C., et al. 1996. Persistent LacZ expression in skeletal muscle of immunodeficient (SCID) mice mediated by highly defective herpes simplex virus type 1 vector. Hum. Gene Ther. 8:439–452. 58. Reed, Clark K., Sferra T. J., Johnson P. R. 1997. Recombinant adeno-associated virus vectors mediate long-term transgene expression in muscle. Hum. Gene Ther. 8:659–669. 59. Huard J., Acsadi G., Jani A., et al. 1994. Gene transfer into skeletal muscles by isogenic myoblasts. Hum. Gene Ther. 5:949–958. 60. Floyd S. S., Clemens P. R., Ontell M. R., et al. 1998. Ex vivo gene transfer using adenovirus mediated full-length dystrophin delivery to mature dystrophic muscles. Gene Ther. 5:19–30. 61. Salvatori G., Ferrari G., Messogiorno A., et al. 1993. Retroviral vector–mediated gene transfer into human primary myogenic cells lead to expression in muscle fibers in vivo. Hum. Gene Ther. 4:713–723. 62. Booth D. K., Floyd S. S., Day C. S., et al. 1997. Myoblast mediated ex vivo gene transfer to mature muscle. J. Tissue Eng. 3:125–133.

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63. Qu Z., Balkir L., van Deutekom J. C., Robbins P. D., Pruchnic R., Huard J. 1998. Development of approaches to improve cell survival in myoblast transfer therapy. J. Cell Biol. 142:1257– 1267. 64. Merly F., Huard C., Asselin I., Robbins P. D., Tremblay J. P. 1998. Anti-inflammatory effect of transforming growth factor-beta1 in myoblast transplantation. Transplantation 65:793–799. 65. Kato K., Shimozato O., Hoshi K., Wakimoto H., Hamada H., Yagita H., Okumura K. 1996. Local production of the p40 subunit of interleukin 12 suppresses T-helper 1–mediated immune responses and prevents allogeneic myoblast rejection. Proc. Natl. Acad. Sci. USA 93:9085–9089. 66. Cornetta K., Morgan R. A., Anderson W. F. 1991. Safety issues related to retroviral-mediated gene transfer in humans. Hum. Gene Ther. 2:5–14. 67. Yang N. S., Burkholder J., Roberts B., Martinell B., McCabel D. 1990. In vivo and in vitro gene transfer to mammalian somatic cells by particle bombardment. Proc. Natl. Acad. Sci. USA 87:9568–9572. 68. Chen C., Okayama H. 1987. High-efficiency transformation of mammalian cells by plasmid DNA. Mol. Cell Biol. 7:2745–2752. 69. Dodds E., Dunckley M. G., Naujoks K., Michaelis U., Dickson G. 1998. Lipofection of cultured mouse muscle cells: a direct comparison of Lipofectamine and DOSPER. Gene Ther. 5:542–551. 70. Trivedi R. A., Dickson G. 1995. Liposome-mediated gene transfer into normal and dystrophindeficient mouse myoblasts. J. Neurochem. 64:2230–2238. 71. Vitiello L., Bockhold K., Joshi P. B., Worton R. G. 1998. Transfection of cultured myoblasts in high serum concentration with DODAC:DOPE liposomes. Gene Ther. 5:1306–1313. 72. Vitiello L., Chonn A., Wasserman J. D., Duff C., Worton R. G. 1996. Condensation of plasmid DNA with polylysine improves liposome-mediated gene transfer into established and primary muscle cells. Gene Ther. 3:396–404. 73. Espinos E., Liu J.-H., Bader C. R., Bernheim L. 2001. Efficient non-viral DNA-mediated gene transfer to human primary myoblasts using electroporation. Neuromusc. Disorders (in press). 74. Lattermann C., Baltzer A. W. A., Whalen J. D., et al. 1998. Gene therapy in sports medicine. Sports Med. Arthrosc. Rev. 6:83–88. 75. Day C. S., Moreland M., Floyd S., Huard J. 1997. Limb-lengthening promotes muscle growth. J. Orthop. Res. 15:227–234.

26 Tissue Engineered Skin Ysabel M. Bello, Anna F. Falabella, and Robert S. Kirsner University of Miami Schosol of Medicine, Miami, Florida

I INTRODUCTION Transfer of autologous skin to close large or refractory wounds or to improve cosmesis has been undertaken for thousands of years. However, limitations of donor site availability, the recurrent nature of chronic wounds, and associated morbidity (slow healing, scarring, pain and infection at the donor site) have led to the use of other skin sources such as allogeneic (derived from another human donor) and/or xenogeneic (derived from another species) skin. These skin sources also are limited by potential problems, including ethical considerations, graft rejection, scarcity of material, risk of infection, and expense of storage. Limitations of these naturally occurring skin replacements have driven the search for synthetic or biosynthetic alternatives. Tissue engineered skin refers to skin products made of cells, extracellular matrix materials, or a combination of cells and matrices [1]. These materials may be derived from a variety of sources such as autologous, allogeneic, or xenogeneic tissue or may be from synthetic materials. II CELLS Critical to the development of skin replacement using living cells was the ability to stratify keratinocytes in culture. In 1975, Rheinwald and Green [2] developed a method that made it possible to grow human keratinocytes serially in vitro. This technique permitted significant expansion of keratinocytes cultures, so that from a 1- to 2-cm2 skin biopsy, up to 1 m2 of a keratinocyte sheet can be generated in about 3 weeks. These cultured cells stratify to the extent that the epidermal graft consists of four to five layers of flattened keratinocytes with a distinct basal layer. Langerhans cells (the professional immune cell of the epidermis) 515

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are absent in the epidermal grafts as are typically melanocytes, although rarely melanocytes may survive the culture process. Using this technique, cultured epidermal autografts were derived. Although limited by fragility and the length of time to grow these sheets, these grafts could “take” and might act as a permanent graft. Initially used for burn patients, patients with other acute and chronic wounds have also been treated. Epicel™ cultured epithelial autografts (Genzyme Tissue Repair, Cambridge, MA) is one of the commercially available epidermal autografts. In a study of 19 burn patients treated with 31 cultured epithelial autografts, the authors reported good adherence in patients with burns involving less than 50% of the total body surface [3]. Using a similar technique with allogeneic cells, a readily available source of skin can be obtained. Cultured epidermal allografts are obtained from allogeneic tissue such as newborn foreskin. However, cultured epidermal allografts as foreign material do not persist, but rather stimulate healing through granulation tissue formation and epithelialization from wound edges. Some have suggested that this is most likely mediated through growthpromoting cytokines. Cultured epidermal allografts can be cryopreserved and stored and then subsequently thawed at room temperature for use. Limited by their fragile nature, they are not currently commercially available. Cultured epidermal allografts have been used to treat burn patients [4], as well as patients with leg ulcers [5]. Although not currently in widespread use, epidermal grafts (auto- and allografts) signify the opportunity to replace skin with epidermal sheets. III MATRICES Due to the fragility of epidermal grafts and the need for dermal replacements in deeper wounds, dermal replacements were developed. These have included acellular and cellular matrices. Alloderm® (LifeCell Corp., Branchburg, NJ) is an acellular collagen matrix, obtained by chemical processing of human cadaver skin to remove the antigenic epidermal and mesenchymal cells, which are targets for immune response. The composition of the dermal matrix and basement membrane is conserved and functions as a template for dermal regeneration. Alloderm has been Food and Drug Administration (FDA) approved for the treatment of burns since 1992 and also has been approved for use in periodontal, plastic, and reconstructive surgery for the filling of subcutaneous tissue defects since 1994. It is cryopreserved and must be reconstituted with normal saline prior to use. Although a multicenter study treating burn patients reported similar outcomes between the use of thin split thickness autografts plus Alloderm versus thicker split thickness skin autografts, Alloderm did allow for harvesting of thinner skin autografts, which resulted in less pain and faster healing rates at donor sites [6]. Integra® (Integra Life Sciences Corp., Plainsboro, New Jersey) is a dermal matrix developed by Burke and Yannas [7]. It is composed of a three-dimensional porous matrix of cross-linked bovine tendon collagen and shark-derived glycosaminoglycans designed to simulate normal human dermis. It is covered by a silastic membrane and serves as a temporary, biosynthetic, bilaminate skin substitute. The FDA approved its use for the treatment of burn wounds in 1996. A controlled, multicenter study of 149 patients with matched-pair thermal injuries was performed to compare the safety and efficacy of Integra versus conventional wound cover (autografts and nonautologous skin such as allograft, xenograft, or dressings) [8]. Integra is used in a two-step process; 3 weeks after initial placement on a

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sterile wound, the silastic membrane is removed and a split thickness skin graft may be placed. Although no significant difference in initial take between allograft and Integra was noted, donor sites healed faster on average in the Integra group 10.6 days compared to 14.3 days for the control site. The advantage of using this matrix includes the ability to take thinner autografts, which allows for more rapid healing of autograft donor sites due to their thinness. An additional finding was that Integra-treated areas had less hypertrophic scarring. Also from a practical standpoint, the silicone layer is also more easily removed than cadaver skin allografts. The FDA requires clinicians to undergo a company-sponsored training program for the use of Integra. IV MATRICES WITH CELLS Transcyte™ (Advanced Tissue Sciences Inc., La Jolla, CA) is a dermal replacement that consists of a nonviable cellular matrix. It is a biosynthetic, bilaminate tissue engineered skin, composed of neonatal human fibroblasts cultured in a three-dimensional nylon mesh. This nylon mesh is coated with porcine dermal collagen. The tissue is covered by a silicone layer. Living fibroblasts synthesize collagen, growth factors, and other extracellular matrix components. Subsequently during processing, the cells are rendered nonviable by freezing. Formerly called Dermagraft TC (transitional covering), Transcyte provides a temporary covering to help protect wounds from fluid loss and to reduce the risk of infection. It is FDA approved for temporary covering of third-degree burns (March 1997) and second-degree burns (October 1997). Transcyte is often used in a two-step process, where after initial placement it is later replaced with a split thickness skin graft. A prospective, randomized, match-controlled trial performed in 14 patients compared the use of Transcyte versus silver sulfadiazine in the treatment of second-degree burns. The authors found that the mean time for 90% epithelialization of wounds was 11 days for Transcyte compared to 18 days for silver sulfadiazine [9]. In this study the Transcyte was sloughed during healing. Another prospective multicenter, randomized, controlled trial performed with 66 patients compared Transcyte with cadaver skin allograft for temporary closure of excised burn wounds. Transcyte was easier to remove and resulted in less bleeding than cadaver skin allograft [10]. Transcyte has a long shelf-life, is cryopreserved, and is immediately available for use. Dermagraft® (Advanced Tissue Science Inc.) is a biosynthetic, dermal matrix with viable living cells. Fibroblasts from neonatal foreskin are cultured into a three-dimensional polyglactin mesh (absorbable suture material). During cell growth, extracellular proteins, collagen and growth factors are released and deposited within this dermis. It can be cryopreserved at 70°C for up to 2 years, however once it is thawed it should be used immediately. Dermagraft is currently available in Canada and Europe, and it has been recently approved by the FDA in the U.S. for the treatment of diabetic foot ulcers. A larger multicenter prospective, randomized, single-blind, controlled trial that enrolled 281 patients with diabetic foot ulcers evaluated the efficacy of Dermagraft application once per week for 8 weeks [11]. The initial results showed that healing was the same between the Dermagraft group and the control group (38.5% versus 31.7%). However, it was found that only 61 patients in the Dermagraft group received tissue that had a high index of metabolic activity. In this subgroup of patients, complete healing was reported 50.8% at 12 weeks [11]. A subsequent large multicenter study undertaken to confirm this finding showed that 30% of patients treated with dermagraft compared to 18% of control patients. Apligraf™ (Graftskin) is a living skin equivalent originally developed by Bell et al. in 1981 [12]. It is now manufactured by Organogenesis (Canton, MA) and marketed by No-

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vartis (East Hanover, NJ). Apligraf contains living allogeneic human neonatal foreskin keratinocytes and living allogeneic neonatal foreskin fibroblasts that are cultured on bovine type I collagen matrix. Apligraf is a biosynthetic, bilaminate skin equivalent (Fig. 1). It was the first biomedical device containing living human cells to be approved by the FDA. First approved in May 1998 for the treatment of venous ulcers, it was FDA approved for the treatment of neuropathic diabetic ulcers in June 2000. As a living bilayered product, it is very similar to human skin, but does not contain Langerhan’s cells, melanocytes, lymphocytes, or endothelial cells. Apligraf has a 5-day shelf-life. Apligraf was studied in a multicenter, prospective, randomized trial involving 293 patients with venous leg ulcer, which compared Apligraf plus compression therapy versus compression therapy alone [13]. Apligraf plus compression was superior to compression alone (63% versus 49%) and the patients healed faster (63 days versus 181 days) (Fig. 2). The greatest difference was seen in ulcers present longer than 6 months. Another multicenter, prospective, randomized, controlled trial involving 208 patients with diabetic foot ulcer compared Apligraf plus conventional therapy with conventional therapy [14]. Apligraf was applied up to five times over a 1-month period, 112 patients received Apligraf and 96 were in the control group. At 12 weeks, Apligraf-treated patients healed to a greatest extent (56% versus 38%). Additionally the average time for 100% wound closure in the Apligraf-treated group was significantly less than the control group (65 days versus 90 days). Apligraf has been used in the treatment of acute wounds as well. A single-center, randomized paired comparison study in 20 patients with three donor site wounds compared healing time, pain, and cosmetic outcomes between Apligraf, split thickness skin autograft, and occlusive dressing (film) [15]. Apligraf was comparable to autograft and superior to film dressing. In a large

Figure 1 Apligraf can be meshed or fenestrated. (A) It can be meshed to a 1.5-to-1 ratio using a mesher. (B) Meshed Apligraf. (C) It can be fenestrated using a scalpel to made the slits. (D) Fenestrated Apligraf.

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Figure 2 (A) Nonhealing venous ulcer. (B) Venous ulcer covered with Apligraf. (C) Venous ulcer healed after treatment with Apligraf plus compression therapy.

series involving surgical excision wounds, Apligraf was found to be safe and effective [16]. Apligraf has been used to treat patients with epidermolysis bullosa. A single-center, prospective open trial performed in 15 patients with 69 acute wounds and nine chronic wounds reported that 75% of the acute wounds were healed at 6 weeks post-treatment, and 75% remained healed at 18 weeks. With chronic wounds, four of nine were healed at 6 weeks, but none of four that were followed up at 18 weeks remained healed.[17] Cultured skin substitute was developed by Hansborough and Boyce [18]; it is a bilayered skin equivalent; it contains autologous fibroblasts growing on a three-dimensional matrix of bovine collagen and glycosaminoglycan and autologous keratinocytes. It is currently under investigation and is not commercially available. Composite cultured skin (CCS or OrCel) (Ortec International Corp., New York, NY) is a bilayered skin equivalent composed of a collagen sponge seeded with allogeneic fibroblasts and keratinocytes cultured from neonatal foreskin. The sponge is coated on one surface with nonporous collagen. Fibroblasts are seeded on the porous aspect of the sponge, while keratinocytes are seeded on the nonporous collagen. It is FDA approved for mitten hand deformities in patients with epidermolysis bullosa and for donor site wounds in burn patients. In this latter group, OrCel treated wounds speed healing and reduce time to re-harvesting compared to the control biobrane treated sites. Clinical trials are underway for venous leg ulcers and diabetic foot ulcers. This product can be cryopreserved. It is predicted that the shelf-life will be about 18 months. Endothelialized skin equivalent (ESE) consists of an in vitro collagen glycosaminoglycan sponge cocultured with keratinocytes, dermal fibroblasts, and umbilical vein endothelial cells. Keratinocytes and fibroblasts were isolated from human skin breast tissue after breast reduction surgery and umbilical vein endothelial cells from healthy human newborns. It is the first endothelialized human tissue-engineered skin in which a network of capillary-like tubes is formed [19]. Using ESE may accelerate graft revascularization by inosculation of its preexisting capillary-like network with the patient’s own blood vessels. However, endothelial cells are quite antigenic, and use of allogeneic cells may lead to rejection.

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V CONCLUSION The development of bioengineered skin has had a significant impact on the treatment of both acute and chronic wounds. It has enabled clinicians to harness the normal healing capacity of skin and use it therapeutically. In addition to treating acute and chronic wounds, tissue engineered skin may also have the capacity to replace diseased skin or be genetically engineered to selectively produce growth factors and other proteins. REFERENCES 1. Eaglstein W. H., Falanga V. 1998. Tissue engineering for skin: an update. J. Am. Acad. Dermatol. 39:1007–1010. 2. Rheinwald J., Green H. 1975. Serial cultivation of strains of human epidermal keratinocytes: formation of keratinizing colonies from single cells. Cell 6:331–344. 3. Rue L. I., Cloffi W., McManus W., et al. 1993. Wound closure and outcome in extensively burned patients treated with cultured autologous keratinocytes. J. Trauma 34:662–667. 4. De Luca M., et al. 1989. Multicenter experience in the treatment of burns with autologous and allogeneic cultured epithelium, fresh or preserved in a frozen state. Burns 15:303–309. 5. De Luca M., Albanese E., Cancedda R., et al. 1992. Treatment of leg ulcers with cryopreserved allogenic cultured epithelium. Arch. Dermatol. 128:633–638. 6. Wainwright D., Malden M., Luterman A., et al. 1996. Clinical evaluation of an acellular allograft dermal matrix in full-thickness burns. J. Burn Care Rehabil. 17:124–136. 7. Burke J., Yannas I., Quinby W., Bondoc C., Jung W., 1981. Successful use of physiologically acceptable artificial skin in the treatment of extensive burn injury. Ann. Surg. 194:413–428. 8. Heimbach D., Luterman A., Burke J., et al. 1988. Artificial dermis for major burns. A multicenter randomized clinical trial. Ann. Surg. 208(3):313–320. 9. Noordenbos J., et al. 1999. Safety and efficacy of Transcyte for the treatment of partial thickness burns. J. Burn Care Rehabil. 20:275–278. 10. Purdue G. F., Hunt J. L., Still Jr. J. M., et al. 1997. A multicenter clinical trial of a biosynthetic skin replacement, Dermagraft-TC, compared with cryopreserved human cadaver skin for temporary coverage of excised burn wounds. J. Burn Care Rehabil. 18(Part 1):52–57. 11. Pollak R., et al. 1997. A human dermal replacement for diabetic foot ulcers. Wounds 9:175–183. 12. Bell E., Ehrlich H., Buttle D., Nakatsuji T. 1981. Living tissue formed in vitro and accepted as skin-equivalent tissue of full thickness. Science 211:1052–1054. 13. Falanga V., et al. 1998. Rapid healing of venous ulcers and lack of clinical rejection. Human skin equivalent investigators group [see comments]. Arch. Dermatol. 134(3):293–300. 14. Veves A., Falanga V., Armstrong D. G., Sabolinski M. L. 2001. Graftskin, a human skin equivalent, is effective in the management of noninfected neuropathic diabetic foot ulcers. Diabetes Care 2001 Feb 24(2):290–295. 15. Muhart M., McFalls S., Kirsner R. S., et al. 1999. Behavior of tissue-engineered skin: a comparison of a living skin equivalent, autograft, and occlusive dressing in human donor sites [see comments]. Arch. Dermatol. 135(8):913–918. 16. Eaglstein W. H., Alvarez O. M., Auletta M., et al. 1999. Acute excisional wounds treated with a tissue-engineered skin (Apligraf). Dermatol. Surg. 25(3):195–201. 17. Falabella A. F., Valencia I. C., Eaglstein W. H., Schachner L. A. 2000. Tissue-engineered skin (Apligraf) in the healing of patients with epidermolysis bullosa wounds. Arch. Dermatol. 136(10):1225–1230. 18. Hansbrough J., Boyce S., Cooper M., Foreman T. 1989. Burn wound closure with cultured autologous keratinocytes and fibroblasts attached to a collagen–glycosaminoglycan substrate. JAMA 262:2125–2130. 19. Black A. F., et al. 1998. In vitro reconstruction of a human capillary-like network in a tissueengineered skin equivalent. FASEB J. 12(13): 1331–1340.

27 The Use of Muscle-Derived Cells for the Treatment of Muscle Pathologies Daniel Skuk and Jacques P. Tremblay Centre de Recherche du Centre Hospitalier de l’Université Laval, Ste.-Foy, Québec, Canada

I INTRODUCTION Tissue engineering aims to restore, maintain, or improve tissue functions that are defective or have been lost by different pathological conditions, either by developing biological substitutes or by reconstructing tissues. The general strategies adopted by tissue engineering can be classified into three groups [1]: (1) implantation of isolated cells or cell substitutes into the organism, (2) delivering of tissue-inducing substances (such as growth factors), and (3) placing cells on or within different matrices. The last of these strategies is more frequently associated with the concept of tissue engineering, i.e., the use of living cells seeded on a natural or synthetic extracellular substrate to create implantable “pieces” of the organism [2]. Influenced by this principle, muscle tissue engineering has been defined as a discipline aimed toward “building in vitro replacements for in vivo problems” [3]. This implies seeding cultured myogenic cells on natural or synthetic substrates to form implantable analogs of muscle tissue able to replace the muscle lost by trauma, surgery, or congenital malformation [4]. So far there have been only a few studies on this subject and their results were limited [4–8]. More extensively studied was the intramuscular implantation of myogenic cells, usually called “myoblast transplantation” or, less frequently, “myoblast transfer therapy.” In this case, the strategy is comprised in the first group indicated, i.e., the implantation of isolated cells in the organism [1]. Instead of prefabricating the muscle tissue in vitro, myoblast transplantation profits of the preexisting tissue structure for an efficient incorporation of the implanted cells in the host. As shown in this chapter, myoblast transplantation success will be conditioned by factors that must be quantitatively determined, such as percentage of 521

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muscle fibers that must express the reporter gene, number of cells to be delivered by distance of injection, distance and topography of injections, etc. These quantitative aspects of myoblast transplantation should reinforce its quality of an “engineering” discipline. To differentiate the two approaches of muscle tissue engineering, avoiding definitions that can be exclusive, we propose to refer to myoblast transplantation as a technique of “in vivo muscle tissue engineering” and the creation of implantable muscle analogs by seeding myogenic cells on matrices as “in vitro muscle tissue engineering.”

II BASIS OF MYOBLAST TRANSPLANTATION The aim of intramuscular implantation of donor myogenic cells is to treat pathological processes of the skeletal muscle which impair their function producing loss of force and movement. Most of the work on myoblast transplantation has been made to develop a treatment for Duchenne muscular dystrophy (DMD), which became the prototype of the muscle diseases to be targeted by this therapeutic approach. At the molecular level, DMD is caused by the genetic absence of dystrophin, a subsarcolemmal fibrilar protein that connects actin in the external sarcomers with a complex of proteins into or in relation with the sarcolemma (for a review see Ref. 9). The absence of dystrophin causes progressive muscle degeneration, which is expressed clinically by rapid loss of muscle force early in the childhood, severe weakness by the age of 10, and death by the age of 20. It was this clinical severity, in addition to its relative frequency, which placed DMD as the principal target of myoblast transplantation, although other muscle pathologies may also benefit from this technique [10–12], including heart diseases [13,14]. Three degrees of therapeutic benefits must be the objective of myoblast transplantation. In a sequence of increasing clinical improvement, they are 1. To slow down the process of muscle degeneration and force loss 2. To stop the pathological process 3. To increase the force in wasted muscles Notably, the first two degrees of clinical improvement would be beneficial for patients at the first stages of a myopathy. The more the patient is affected, the more will it be necessary to reach the third objective. For these therapeutic purposes, myoblast transplantation may act by two complementary actions: 1.

2.

Acting as a vehicle for delivery of normal genes. Normal donor myoblasts induce “gene complementation” in the host myofibers, i.e., the normal donor nuclei incorporated by the host myofibers will allow the expression of the protein which is mutated in the host genome. In the specific case of DMD, myoblast transplantation aims to induce the expression of dystrophin into the largest number of myofibers as possible. Those myofibers where donor nuclei coexist with host nuclei are called “hybrid” myofibers. Increasing the myogenic capacity of the muscle. The incorporation of donor nuclei in the host myofibers should allow the synthesis of more contractile proteins. The force generated by a muscle is a function of its total number of contractile proteins and, therefore, of the number and diameter of its myofibers. In the case of DMD, the progression of the disease causes the irreversible and progressive reduction in the number and diameter of myofibers. This myofiber loss is at-

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tributed to the fact that successive cycles of myofiber necrosis and regeneration decimate the capacity of satellite cell