 |
 |
 |
|
| home | CME | help | feedback | subscriptions | archive |
|
© 2004 The Journal of Bone and Joint Surgery, Inc.
Engineering Principles of Clinical Cell-Based Tissue Engineering
1 Departments of Orthopaedic Surgery and Biomedical Engineering (A41), The
Cleveland Clinic Foundation, 9500 Euclid Avenue, Cleveland, OH 44195. E-mail
address for G.F. Muschler:
muschlg{at}ccf.org
Investigation performed at the Departments of Orthopaedic Surgery and
Biomedical Engineering, The Cleveland Clinic Foundation, Cleveland, Ohio, and
the Departments of Biological Engineering and Mechanical Engineering,
Massachusetts Institute of Technology, Cambridge, Massachusetts
Tissue engineering is a rapidly evolving discipline that seeks to repair, replace, or regenerate specific tissues or organs by translating fundamental knowledge in physics, chemistry, and biology into practical and effective materials, devices, systems, and clinical strategies. Stem cells and progenitors that are capable of forming new tissue with one or more connective tissue phenotypes are available from many adult tissues and are defined as connective tissue progenitors. There are four major cell-based tissue-engineering strategies: (1) targeting local connective tissue progenitors where new tissue is desired, (2) transplanting autogenous connective tissue progenitors, (3) transplanting culture-expanded or modified connective tissue progenitors, and (4) transplanting fully formed tissue generated in vitro or in vivo. Stem cell function is controlled by changes in stem cell activation and self-renewal or by changes in the proliferation, migration, differentiation, or survival of the progeny of stem cell activation, the downstream progenitor cells. Three-dimensional porous scaffolds promote new tissue formation by providing a surface and void volume that promotes the attachment, migration, proliferation, and desired differentiation of connective tissue progenitors throughout the region where new tissue is needed. Critical variables in scaffold design and function include the bulk material or materials from which it is made, the three-dimensional architecture, the surface chemistry, the mechanical properties, the initial environment in the area of the scaffold, and the late scaffold environment, which is often determined by degradation characteristics. Local presentation or delivery of bioactive molecules can change the function of connective tissue progenitors (activation, proliferation, migration, differentiation, or survival) in a manner that results in new or enhanced local tissue formation. All cells require access to substrate molecules (oxygen, glucose, and amino acids). A balance between consumption and local delivery of these substrates is needed if cells are to survive. Transplanted cells are particularly vulnerable. Theoretical calculations can be used to explore the relationships among cell density, diffusion distance, and cell viability within a graft and to design improved strategies for transplantation of connective tissue progenitors. Rational strategies for tissue engineering seek to optimize new tissue formation through the logical selection of conditions that modulate the performance of connective tissue progenitors in a graft site to produce a desired tissue. This increasingly involves strategies that combine cells, matrices, inductive stimuli, and techniques that enhance the survival and performance of local or transplanted connective tissue progenitors.
The discipline of engineering translates fundamental knowledge in physics, chemistry, and biology into materials, devices, systems, and strategies to achieve practical benefits. This discipline also includes the systematic definition and assessment of each variable that may contribute to the success or failure of any engineering effort. Tissue engineering applies this conceptual framework to advance the repair, replacement, or regeneration of organs and tissues1-3. Current tissue-engineering strategies include transplantation of whole organs or tissues with use of pedicle flaps and microvascular techniques, transplantation of thin sections of tissues (e.g., split-thickness skin grafts), transplantation of cell suspensions (e.g., blood transfusions or bone marrow transplants), and endoprosthetic replacement of tissues. In orthopaedics, bone4, cartilage5,6, tendon7,8, ligament9, meniscus, intervertebral disc10,11, fat, muscle12, and nerve are the primary targets. In recent years, the options for orthopaedic tissue engineering have increased dramatically. These options include methods for harvest and transplantation of tissue-forming cells, the use of an expanding array of bioactive matrix materials as tissue scaffolds, local or systemic delivery of commercially available peptide hormones and growth factors, and other methods to control the local chemical and biophysical environment. These new options highlight a transition from the historically materials-based tissue-level approach, with which mechanically durable, bioinert, or biocompatible materials were preferred, to a focus on cell-based or bioactive materials and stimuli. This evolving approach focuses on the function of cells and the role of materials, implants, and biophysical stimuli in modulating cell function. Cell-based tissue-engineering tools and methods create exciting new opportunities that might be useful in a broad array of potential clinical applications. These opportunities also precipitate a critical need for orthopaedic surgeons to participate actively in the design, development, critical evaluation, and informed use of these methods. Active participation requires that orthopaedic surgeons have a solid foundation in the contemporary concepts and principles of cell-based tissue engineering. This article reviews the central paradigms of contemporary tissue engineering. Specifically, it addresses stem cells and progenitor cells in musculoskeletal tissues (the cells responsible for all new tissue formation), strategies for the clinical use of these cells, barriers to cell transplantation and cell survival, and strategies and variables in the design and optimization of cell-based tissue-engineering scaffolds.
Stem cells and progenitor cells are present in all adult tissues and are critical to tissue health, maintenance, and response to injury or disease throughout life. Stem cells are the source of all new tissues arising from repair and remodeling and are modulated by chemical and physical signals that control their activation, proliferation, migration, differentiation, and survival. Stem cells give rise to progenitor cells and are distinguished from progenitor cells by their capacity for self-renewal, or self-regeneration, by a process of asymmetric cell division (Fig. 1). In contrast, progenitor cells (also called transit cells) proliferate and expand in number. Progenitor cells have a limited capacity for self-renewal and are committed to progress toward a different or more differentiated phenotype (Fig. 2)13,14.
Stem cells and progenitor cell populations are part of continuous systems involving cell loss and regeneration in virtually all human tissues that consist of ongoing generation of new cells and the orderly transition of cells from one state to another. This turnover is most evident in the lining cells of the gastrointestinal tract (every three days) or dermis (every fourteen days). In musculoskeletal tissues, turnover is much slower and has been best characterized in bone, where the cells that give rise to and support bone tissue progress through a series of stages beginning upstream with the stem cell. Stem cells give rise to progenitor cells, which progress to become pre-osteoblasts and then osteoblasts. Osteoblasts represent yet another transit population, with a life span of only about forty days, and give rise to both the matrix of new bone tissue and the downstream cells that comprise bone tissue (i.e., osteocytes and bone-lining cells). As osteoblasts reach the end of their functional life, they have three possible fates: they may become osteocytes, they may become lining cells on the surface of mature bone, or they may die by means of apoptosis. As an osteocyte or a lining cell, the same cell may survive for a mean of twenty years or more in human cortical bone, until the region of bone in which it resides is remodeled by the progeny of yet another wave of stem cell function14. Bone repair and regeneration following a fracture or a bone-grafting procedure follow the same steps. The same principles apply to cells in muscle, tendon, ligament, and cartilage15. An intriguing feature of many tissues, including musculoskeletal tissues, is that upstream progenitors are often multipotent. Stem cells derived from bone, bone marrow, muscle, and fat have all been shown to be capable of differentiation into multiple phenotypes, including bone, cartilage, tendon, ligament, fat, muscle, and nerve16-20. This has important implications with regard to the design of tissue engineering strategies, in that cells derived from one tissue might be useful in forming a different tissue. Sources of musculoskeletal stem cells and progenitor cells include bone marrow, peritrabecular tissues in cancellous bone, periosteum, cartilage, muscle, fat, and vascular pericytes20-27. The harvest of these tissues varies with respect to the associated host morbidity28-34. Aspiration of bone marrow (Figs. 3-A and 3-B) is associated with the least morbidity and provides a single cell suspension that can be readily processed intraoperatively for immediate implantation24. Fat has also been proposed as a low-morbidity tissue source, although it requires greater processing20,35.
The number of stem cells and progenitor cells in various tissues can be assayed in vitro by liberating the cells from the tissues and growing them in tissue culture under conditions that promote activation and proliferation of stem cells and progenitor cells. The number of stem cells and progenitors can be estimated on the basis of the number of colony-forming units16,19 (Fig. 4). The term connective tissue progenitors has been used to denote this combined heterogeneous population of stem cells and progenitor cells that is capable of both proliferation (colony formation) and differentiation into one or more connective tissue phenotypes13.
Cells derived from connective tissue progenitors can be expanded (grown) in vitro for use in research or for tissue engineering applications. Under these conditions, isolated cells rapidly become more uniform. When cells are grown in vitro, clones of cells that divide most rapidly and those that have the greatest capacity for continued proliferation have a competitive advantage. In vitro expansion therefore produces a selective pressure favoring these traits. Culture-expanded and selected cell populations have been ascribed various names, including bone marrow stromal cells16, mesenchymal stem cells17, and adult multipotential progenitor cells36,37. These terms are not synonymous with regard to their precise definitions and biological capabilities13. However, all denote that progenitor cells can be isolated and expanded under appropriate conditions and that these cells can retain the capacity to differentiate into a variety of musculoskeletal phenotypes16-18. All tissues vary substantially with respect to cellularity and the prevalence of connective tissue progenitors. Aspirated bone marrow is the best characterized source of connective tissue progenitors, containing a mean of approximately forty million nucleated cells and approximately 2000 connective tissue progenitors per milliliter, or about one connective tissue progenitor per 20,000 cells, when an appropriate technique is used24,33,38. In contrast, fat and muscle tissues are far less cellular (approximately six million cells per cubic centimeter of tissue), but the prevalence of connective tissue progenitors in fat (as high as one per 4000 cells) may be slightly higher than that in bone marrow. Differences between connective tissue progenitors harvested from various tissue sources and individuals are only beginning to be understood. These differences depend on the health and histological characteristics of the local tissues and the kinetics of the stem cell function at the site. These variables are in turn influenced by age, gender, and both local and systemic disease13,24,38,39. For example, bone marrow cellularity has been shown to decline with age. There is also an age-related decline in the prevalence of connective tissue progenitors, at least in women38. However, age and gender account for only a small fraction of the variation in the concentration and prevalence of connective tissue progenitors between patients24,33,38. As a result, osteogenic connective tissue progenitors can be harvested by bone marrow aspiration in patients of all ages. Differences in biological potential among connective tissue progenitors derived from various tissues can have important practical implications with regard to the selection of cell sources for tissue engineering. Bone-marrow-derived connective tissue progenitors include cells that are capable of differentiation into a broad range of phenotypes, including bone, fibrous tissue, fat, muscle, cartilage, and perhaps even neural tissue, liver, and cardiac muscle13,14,16,17,40-42. In contrast, connective tissue progenitors from adult articular cartilage are only capable of forming cartilage. Connective tissue progenitors derived from muscle and fat may also have a broad intrinsic differentiation repertoire20,27,30-32. Some studies have suggested that fat-derived and bone-marrow-derived cells are similar43, but others have demonstrated a decreased osteogenic potential in fat-derived cells44 and the absence of surface markers characteristic of osteoblastic progenitors45.
There are four major types of cell-based tissue engineering: (1) local targeting of connective tissue progenitors where new tissue is needed, (2) transplanting autogenous connective tissue progenitors to augment the local population, (3) transplanting culture-expanded or modified connective tissue progenitors, and (4) transplanting fully formed tissue.
Targeting Connective Tissue Progenitors In Situ Locally delivered growth factors (e.g., bone morphogenetic proteins [BMPs], fibroblast growth factor-2 [FGF-2], and vascular endothelial growth factor [VEGF]) also target local cells. The capacity of some growth factors to selectively activate bone-forming connective tissue progenitors and/or enhance the probability that their progeny will differentiate into bone has been defined as osteoinduction. Again, the concept of induction can apply equally well to stimuli that promote activation and differentiation of connective tissue progenitors toward any desired phenotype. Biophysical stimulation, such as mechanical loading48-50, electromagnetic stimulation51,52, or ultrasound53, is also an example of cell targeting. Systemic pharmacological strategies, such as the use of parathyroid hormone for the treatment of osteoporosis54-59 or the use of systemic growth hormones to induce an increase in muscle mass in the elderly60,61, are types of cell-targeted tissue engineering as well.
Transplantation of Autogenous Connective Tissue Progenitors Autogenous cancellous bone-grafting has long been the most prevalent and relatively effective example of cell transplantation, although only a small fraction of the transplanted cells actually survive63-65. In the past decade, several additional transplantation strategies have been introduced. Several uncontrolled clinical studies have suggested that transplantation of connective tissue progenitors in aspirated bone marrow has value in bone-healing applications22,66,67. In an uncontrolled, nonrandomized consecutive study, Connolly et al.68,69 found bone marrow injection to be successful in the treatment of eighteen of twenty tibial nonunions. Connolly et al.25 also showed that concentration of bone marrow cells with use of a centrifuge could increase osteogenesis further, a strategy that has been supported by other investigators70,71. Many surgeons now use bone marrow because of its biological value and low risk. One of us (G.F.M.) has had clinical experience (albeit without independent or detailed prospective documentation) with aspiration of bone marrow in more than 900 patients undergoing elective orthopaedic procedures over the past fifteen years. The aspiration volumes in the patients ranged from 16 to 200 mL. Patient cohorts representing a subset of this experience have been reported on in two publications24,33. There were only two reported bruises, no hematomas, no infections, and no chronic pain at the aspiration site. On direct questioning, most patients reported no pain at the aspiration site during their hospitalization, and in no case was the bone marrow aspiration site the reason for the patient taking pain medication, a factor limiting rehabilitation, or the cause for a delay in discharge from the hospital.
The aspiration technique is important. One of us (G.F.M.) and
colleagues24 found
that limiting the volume of the aspirate to
Transplantation of Culture-Expanded Autogenous Cells
Transplantation of Genetically Modified Cells The biological risk associated with genetic manipulation is greater than that associated with the alternatives. As a result, demonstration of safety is currently as great a challenge as is demonstration of biological efficacy89. While transplantation of genetically modified cells may not play a role in elective clinical tissue engineering in the near future, it has substantial potential value, particularly in the setting of inherited genetic defects (e.g., osteogenesis imperfecta)90 and in tissues, such as cartilage, that consist of relatively homogeneous long-lived cells and in which phenotypic expression may be induced and durably maintained by expression of a single gene86,91.
Ex Vivo Tissue Generation and Transplantation In some areas, transplantation of thin tissue grafts (e.g., cartilage, corneal, and skin grafts) is possible without an immediate connection to a developed vascular system. However, in most tissues, cell survival requires a functioning vascular system. Some reports describe ex vivo generation of vascular transport systems92,93, but these approaches are not currently clinically practical.
In all settings in which cells are transplanted, access to substrate molecules (oxygen, glucose, and amino acids) and clearance of products of metabolism (CO2, lactate, and urea) are critical to cell survival. The movement of these molecules in and out of the graft site is collectively referred to as mass transport. Mass transport can be mediated by fluid flow (convection), both in the circulatory system and in the extracellular space, between the vessel lumen and the cell membrane. The pressure gradients driving this fluid flow can be induced by tissue deformation (movement), mechanical loading, muscle contraction, gravitational pooling, Starling flow, and arterial pulsation. Convection can be particularly important when cells are embedded in a dense extracellular matrix, such as bone and cartilage, where it has been shown to play a major role in enhancing the transport of large molecules (e.g., proteins and growth factors)94-97. However, in most tissues, passive diffusion along concentration gradients is the principal mechanism for mass transport, particularly for small molecules. In metabolically active tissues such as trabecular bone and bone marrow, the distance that oxygen must diffuse between a capillary lumen and a cell membrane is almost never more than 40 to 200 µm98,99. This diffusion distance is critical to maintaining the balance between oxygen delivery to a site and consumption of oxygen by cells, both in native tissues and in tissue engineering strategies involving cell transplantation. When cells are transplanted clinically, the vessels that deliver oxygen are initially confined to the outer surface of the graft site. As one moves deeper into the graft site, each transplanted cell competes for oxygen and other nutrients with other transplanted cells. Transplanted cells also compete with other cells that are recruited as part of the local inflammatory response following implantation. In most clinical grafts, the diffusion distance for oxygen and other metabolites from the edge of the graft to the center of the graft is a minimum of 5 mm, or approximately fifty times the normal diffusion distance. In this setting, diffusion is able to support only a limited number of transplanted cells before the balance between metabolic demand and diffusion creates a zone in the center of the graft where oxygen tension is too low to support viable cells, resulting in central necrosis. The size of the necrotic region and the number, distribution, and type of cells that do survive in the deeper regions of the graft site are a function of many variables, which can be analyzed with use of basic engineering principles. The principal variables are the concentration of oxygen at the surface of the graft site (CO), the concentration and distribution of cells in the site (including inflammatory cells), the rate of oxygen consumption by cells within the site, fluid flow within the site, the diffusion constant for oxygen, and the biological response of cells (survival, proliferation, migration, and differentiation) to hypoxia. Relatively little is known about the balance between diffusion and consumption of proteins, peptides, or the signaling molecules. However, because diffusion of oxygen is relatively slow and oxygen consumption is high, the transport of other nutrients (e.g., glucose and amino acids) is generally more favorable than that of oxygen. Oxygen is therefore the limiting factor in cell survival in most grafts. As a result, few cells tolerate diffusion distances of >0.2 mm. For example, rat osteoblasts seeded on porous scaffolds in vitro form a viable tissue that is no greater than 0.2 mm thick100. Islet cells show necrosis when the diffusion distance exceeds approximately 0.1 mm101-103. Cartilage is exceptional, maintaining viability in avascular regions >1 mm thick104, although oxygen transport in cartilage in vivo may be enhanced by convective flow95,105.
Theoretical modeling can be used to explore the relationships among cell
density, diffusion distance, and cell viability within a graft.
Figures 5-A and 5-B illustrate
a theoretical cellular implant, where diffusion of oxygen in the x-direction
is balanced by cellular consumption. When the differential equation describing
the balance between diffusion and reaction is made nondimensional, a
parameter, denoted
Several factors may modify these calculated estimates. First, because cell delivery systems are generally prepared in room air, the initial implant is usually saturated with oxygen. This dissolved oxygen will support cell respiration for at least several hours after implantation, blunting the abrupt decrease in regional oxygen concentration. Second, not all implanted cells continue to respire at basal rates. Cells that are very sensitive to the trauma of transplantation may die, promptly reducing the initial metabolic demand within the implant. However, cell death also results in the local release of products of cell lysis and adds debris to the site. This may increase the intensity of local inflammation and the metabolic demand from inflammatory cells at a later time. The survival of transplanted connective tissue progenitors also depends on the response of these cells to the transplantation environment. Observation has long supported the concept that at least some connective tissue progenitors in bone and bone marrow have a high capacity to survive in hypoxic conditions63,64,108. Experimental data also have shown that many stem and progenitor cells, including connective tissue progenitors in bone, exhibit a remarkable tolerance to, and are even stimulated by109-113, hypoxia, not unlike endothelial cells114,115. The capacity to convert to glycolysis transiently in response to hypoxia is one adaptive mechanism116. The rate and extent of revascularization are also critical. Prompt revascularization favors osteoblastic differentiation, whereas prolonged hypoxia favors formation of cartilage or fibrous tissue17,117-119. These concepts can be converted into several practical strategies to optimize cell survival in clinical grafts. One method is to reduce the concentration of transplanted cells. Another is to limit the transplanted cells to only those cells that contribute to the formation of the desired tissue (i.e., connective tissue progenitors and perhaps endothelial cells), while excluding red blood cells and the vast majority of other nucleated cells. Both concentration25 and selection strategies62 have been shown to enhance graft performance. Finally, in the future, culture-expanded cells might be preadapted to hypoxic conditions prior to transplantation and/or selected to enrich for those most likely to survive. Other options to improve the local matrix environment and to enhance mass transport are discussed below.
Three-dimensional porous scaffolds play a critical role in both cell targeting and cell transplantation strategies. Scaffold matrices serve as space-holders to prevent encroachment of surrounding tissues into the graft site. They provide surfaces that facilitate the attachment, survival, migration, proliferation, and differentiation of stem cells and progenitors. They also provide a void volume in which vascularization, new tissue formation, and remodeling can occur (Fig. 6). In addition, scaffolds can provide a vehicle for delivery of cells into a graft site, facilitating their retention and distribution throughout the region where new tissue is desired1,46,120-122.
A broad range of scaffolds is already available for clinical use, and many new scaffolds are under development. Differences between scaffolds can generally be categorized into one or more of six domains: bulk material, three-dimensional architecture and porosity, surface chemistry, mechanical properties, initial scaffold environment (osmolarity and pH), and late scaffold environment (degradation characteristics). Each domain has important implications with respect to the biological response to a scaffold and its utility in transplanting or supporting local stem cells and progenitors.
Bulk Materials
Three-Dimensional Architecture and Porosity
Options for the structural design of tissue scaffolds are almost infinite. The macrostructures include regular geometric shapes (e.g., blocks, pellets, and dowels), amorphous structures (e.g., randomly packed chips, granules, or fibers), randomly integrated structures (e.g., foams or freeze-dried materials), and formally designed regular structures (e.g., machined, printed, woven, or assembled structures). Gel or putty preparations can also be made from powders or fibers, by mixing them with plasticizing agents (e.g., glycerol, cellulose, and hyaluronan) or by conducting in situ polymerization with use of chemical, photochemical, or enzymatic methods120. In some cases, a desirable structure has been borrowed from nature, such as the highly interconnected porous structure of cancellous bone or some corals124. Selective processing (machining, size and density selection, washing, and demineralization) now provides a variety of relatively optimized materials for use in special clinical settings. Most methods for fabricating porous scaffolds—i.e., particulate leaching100, freeze-drying125, gas infusion126, and phase separation120—create isotropically distributed voids and connecting pores (such as in a sponge) by using particles or bubbles when the scaffold is solidified. In the past decade, substantial advances have been made in the methods for producing more precise hierarchical microstructures from a variety of materials. These are now being applied to create strategically oriented channels and pores and defined macroscopic shapes. The most notable innovations involve solid free-form fabrication methods (e.g., three-dimensional printing process and stereolithography)1,120, which provide a feature resolution of approximately 200 µm. Creation of more defined porous structures offers the potential for greater control over the distribution of bulk material within a graft site as well as control over patterns of cell migration, fluid flow, and diffusion throughout the device127-129. Nanostructural features (<100 nm) may also play an important role in scaffold function. Nanopores are too small to influence where cells can or cannot migrate, but they may still have important effects on cell behavior by changing surface texture or diffusion of soluble materials. All other features being equal, the presence of interconnected nanoporosity within the walls of a porous structure can open up a much wider path for mass transport, thereby improving cell survival in the scaffold.
Mechanical Properties A scaffold's mechanical properties (strength, modulus, toughness, and ductility) are determined both by the material properties of the bulk material and by its structure (macrostructure, microstructure, and nanostructure). Matching the mechanical properties of a scaffold to the graft environment is critically important so that progression of tissue healing is not limited by mechanical failure of the scaffold prior to successful tissue regeneration. Similarly, because mechanical signals are important mediators of the differentiation of connective tissue progenitors, a scaffold must create an appropriate stress environment throughout the site where new tissue is desired. One of the greatest challenges in scaffold design is the control of the mechanical properties of the scaffold over time. Scaffolds that do not degrade (metals and ceramics) simplify this problem and can provide excellent and durable function in some settings. However, these materials can also compromise tissue repair and function. It is obvious that persistence of a scaffold or implant precludes the formation of new tissue in the space that it occupies. In addition, following integration of a rigid nondegradable implant, adjacent tissue is often mechanically protected (stress-shielded), changing local mechanical signals and resulting in loss of desired local tissue. Stress concentration at the interface between a high-stiffness implant and native tissue can increase the risk of mechanical failure (e.g., fracture) and pain. Finally, if subsequent procedures require removal of the implant scaffold (e.g., because of infection or migration), all of the new tissue within the implant may also be lost, eliminating the value of the initial procedure. Problems arising from retained implants have increased the desire to use resorbable scaffolds whenever feasible. One example of that strategy is the use of impaction grafting for reconstruction of contained periprosthetic defects. Another is the recent shift from the use of very slowly degradable ceramics (e.g., hydroxyapatite) for bone-void fillers to the use of more rapidly resorbed materials (e.g., tricalcium phosphate)130. This demand for resorbable scaffolds continues to fuel the development of resorbable inorganic polymers120. Almost all new materials being developed for tissue engineering ultimately resorb, and three key features of the degradation or resorption process appear to influence performance: the rate at which the matrix loses its mechanical properties, the rate at which the matrix is removed from the site, and the nature and concentration of the soluble products that are released into the site as the material is broken down. Controlled degradation of mechanical properties alone is a major challenge. Mechanical properties can be lost as a result of internal degradation of the bulk material (e.g., hydrolysis) or the accumulation of fatigue damage. In recent years, it has become more evident that, even in controlled settings, the mechanical and chemical degradation of the same polymer can vary substantially between species, individuals, anatomic locations, and clinical settings. As a result, it has been very difficult to define an optimal degradation rate for materials to be used in general clinical practice. In general, most design strategies tend to extend degradation time over months, in order to minimize the risk of early failure in preference to minimizing the risks associated with delayed resorption131.
Surface Chemistry Biological fluids contain a vast diversity of proteins, and cells have hundreds of different types of cell surface receptors. There are twenty-four distinct cell-matrix receptors in the integrin family alone136. As a result, it is not surprising that scaffold materials have been discovered and selected empirically. However, powerful surface analytical techniques are being used to illuminate the protein adsorption properties on various surfaces in an effort to find out why some materials are so favorable for bone cell adhesion and bone formation137. For example, it has been speculated that hydroxyapatite and some other ceramics may preferentially sequester growth factors, growth factor-binding proteins, or adhesion molecules that are important for bone regeneration. Indeed, hydroxyapatite and tricalcium phosphate materials perform successfully as depot delivery vehicles for BMPs, in both animals138 and humans139,140. Like other implanted materials, allograft bone matrices (both mineralized and demineralized) rapidly accumulate biomolecules on their surface, which have biological effects on local cells. However, allograft bone already contains many embedded adhesion molecules and growth factors141. These include the BMPs, although the concentration is far lower than that delivered with use of purified recombinant BMP products and the release is much slower, requiring matrix degradation by local cells. Furthermore, the concentration and presentation of bioactive molecules in allograft bone may vary widely depending on the age, gender, and genetics of the donor; the tissue site of origin; and the tissue processing procedures142. The attachment, survival, proliferation, and differentiation of stem cells and progenitor cells can be modulated in vitro if implants are precoated with selected bioactive proteins143,144. Furthermore, proteins can be selectively concentrated and presented on surfaces with use of nonspecific surface interactions (e.g., dip-coating or lyophilization). This is the strategy that is used to deliver BMP-2 (Infuse; Medtronic Sofamor Danek, Memphis, Tennessee) and OP-1 or BMP-7 (OP-1 Device; Stryker Biotech, Hopkinton, Massachusetts) in two clinical products that are currently available for improving bone-healing145-147. In the case of the BMP-2 product (Infuse), BMP-2 is provided in solution and is dripped into an absorbable collagen sponge. BMP-2 binds to the collagen and is then released. In the case of the OP-1 (BMP-7) product (OP-1 Device), 3500 µg of OP-1 is lyophilized onto 1 g of bovine type-I collagen powder. OP-1 is then released by solubilization from the collagen surface. The pharmacokinetics of delivery of BMPs have been shown to be an important clinical variable in a variety of materials, including degradable polymers148-155, type-I collagen156-158, and calcium phosphate ceramics159,160. The retention time of implanted BMPs have been shown to correlate with biological efficacy, presumably because the longer a BMP is retained, the higher the likelihood that it will act on an appropriate target cell. Protein residence time can be estimated in vivo by measuring the rate of clearance following the implantation of radioactive protein, with the assumption that the protein remains active as long as residual radiation can be measured. Retention has been shown to be related to solubility161 and protein isoelectric point138. The release kinetics of BMP-2 from a degradable collagen sponge in rabbits was described by Bouxsein et al.162. Approximately 25% of the delivered BMP-2 was released rapidly from the implantation site, but as much as 37% remained at the site one week after implantation. The advantage of the current clinical strategies for protein delivery is that they are technically simple. However, they require that the protein be delivered in a high concentration. This delivery strategy presumably allows the protein to diffuse into tissues adjacent to the implant site to act on the local connective tissue progenitor target population. This release pattern may also establish a concentration gradient around the graft site that may be important for chemotactive factors. However, the disadvantage of current strategies is that they provide relatively little control over the rate of delivery, conformation, presentation, clearance, or degradation of the delivered protein. While current strategies for delivery of BMPs can be effective, there is reason to believe, given the very large supraphysiological doses of BMPs that are needed, that the vast majority of the protein that is delivered is wasted and that only a small fraction actually comes into contact with target cells to elicit a receptor-mediated signal that enhances new bone formation. These methods therefore leave substantial room for improvement in the delivery kinetics (rate and duration) and distribution of bioactive proteins to optimize efficiency. Proteins and small bioactive peptides can also be selectively concentrated and presented by covalently linking them to a surface1,120,140. This provides more control over conformation, a slower rate of release from the surface, and longer retention. While this method may not be appropriate for many soluble proteins, particularly signaling molecules that need to be internalized for function (e.g., steroid hormones), there is growing evidence suggesting that presenting growth factors in a matrix-bound fashion not only may be suitable for clinical use, but also may better mimic the native physiology and improve outcomes163-166. This strategy may also be particularly well suited for the design of matrices with selective affinity for specific cells or sets of cells (e.g., connective tissue progenitors, endothelial cells, and platelets).
Initial Scaffold Environment: Osmolarity and pH
Late Scaffold Environment: Degradation Products Polyesters, such as polylactides and polyglycolides, are currently the so-called workhorses of synthetic degradable surgical materials. They have been used for decades as sutures, surgical meshes, and more recently as fixation hardware (e.g., suture anchors and screws)120. The degradation of these materials can be controlled over a range of weeks to years. Polymer hydrophobicity and crystallinity both influence the rate at which water penetrates and hydrolyzes the solid polymer and thus the rate of device breakdown140. However, the degradation products of these materials (lactic acid and glycolic acid) are not ideal for tissue regeneration. Furthermore, they tend to be released as a bolus after a long period of residence. This is due to the fact that they are degraded by hydrolysis in a process that first randomly degrades the bulk polymer, progressively reducing the molecular weight and the mechanical properties of the material but leaving the total mass of polymer essentially the same until the molecular weight of the fragments that are created is small enough to make them soluble. When this occurs, soluble material is generated rapidly, liberating the bulk polymer into solution but creating a profound local decrease in pH. Regression of local bone formation and sterile cysts in bone and soft tissue have been commonly observed15,167. Furthermore, in some settings, the polymer can crystallize as it degrades, creating particles that persist for years. In highly porous devices, the effects of degradation products may be less pronounced because the volume fraction of the bulk polymer in the graft site is smaller and the degradation products that are released are cleared more readily through a broader surface of contact with local extra-cellular fluids and vascular perfusion. Regardless, the use of polylactides and polyglycolides for bone regeneration remains controversial, and devices based on these materials are not widely used in clinical practice. The limitations of polyesters are being addressed by the development of new classes of degradable materials that possess reasonable mechanical strength and do not release acidic degradation products. One material consisting of a copolymer of polyethylene oxide and polybutylene terephthalate showed promising osteoconductive results in animals but failed to induce bone formation in the iliac crests of humans168. A different class of materials, pseudo-polyamino acids, have been synthesized with a range of degradation properties and have been shown to offer improved behavior in bone sites in animals169,170. An innovative approach to matching degradation rate to tissue ingrowth is being developed by Hubbell et al., using polyethylene oxide-based gels that contain both cell adhesion molecules and other peptides, which are selected to provide specific biological activity and are released as the gel is degraded171,172. The gels can be formed in situ, and they function essentially at a synthetic extracellular matrix designed as a true surface eroding polymer. The gels are degraded by specific protea |
Top
Abstract
Introduction
2 mL per site reduces dilution
with peripheral blood and significantly increases the concentration of
marrow-derived connective tissue progenitors (p < 0.001). Recent data have
shown that the efficacy of a bone marrow graft can be significantly enhanced
by the use of the surface of some porous implantable materials to selectively
concentrate and select marrow-derived connective tissue progenitors from bone
marrow (p <
0.001)
2, emerges as the relative rate of reaction
to the relative rate of diffusion. When 
3.3 x 10-17
mol/cell-sec
