Scaffolds provide mechanical support and shape for neotissue construction in vitro and/or through the initial period after implantation as cells expand, differentiate, and organize (Stock and Vacanti, 2001). Materials that mainly have been used to date to formulate degradable scaffolds include synthetic polymers, such as poly(L-lactic acid) (PLLA) and poly(glycolic acid) (PLGA), and polymeric biomaterials, such as alginate, chitosan, collagen, and fibrin (Langer and Tirrell, 2004). Composites of these synthetic or natural polymers with bioactive ceramics such as hydroxyapatite or certain glasses can be designed to yield materials with a range of strengths and porosities, particularly for the engineering of hard tissues (Boccaccini and Blaker, 2005).
A scaffold used for tissue engineering can be considered an artificial extracellular matrix (ECM) (Rosso et al., 2005). It has long been appreciated that the normal biological ECM, in addition to contributing to mechanical integrity, has important signaling and regulatory functions in the development, maintenance, and regeneration of tissues. ECM components, in synergy with soluble signals provided by growth factors and hormones, participate in the tissue-specific control of gene expression through a variety of transduction mechanisms (Blum et al., 1989; Jones et al., 1993; Juliano and Haskill, 1993; Reid et al., 1981). Furthermore, the ECM is itself a dynamic structure that is actively remodeled by the cells with which it interacts (Behonick and Werb, 2003; Birkedal-Hansen, 1995). An important future area of tissue engineering will be to develop improved scaffolds that more nearly recapitulate the biological properties of authentic ECM (Lutolf and Hubbell, 2005).
Decellularized tissues or organs can serve as sources of biological ECM for tissue engineering. The relatively high degree of evolutionary conservation of many ECM components allows the use of xenogeneic materials (often porcine). Various extracellular matrices have been utilized successfully for tissue engineering in animal models, and products incorporating decellularized heart valves, small intestinal submucosa (SIS), and urinary bladder have received regulatory approval for use in human patients (Gilbert et al., 2006). The use of decellularized matrices is likely to expand, because they retain the complex set of molecules and three-dimensional structure of authentic ECM. Despite many advantages, there are also concerns about the use of decel-lularized materials. These include the potential for immu-nogenicity, the possible presence of infectious agents, variability among preparations, and the inability to completely specify and characterize the bioactive components of the material.
Current developments foreshadow the development of a new generation of biomaterials that use defined, purified components to mimic key features of the ECM. Electrospinning allows the production of highly biocompatible micro-and nano-fibrous scaffolds from synthetic materials, such as poly(epsilon-caprolactone), and from diverse matrix proteins, such as collagen, elastin, fibrinogen, and silk fibroin (Boland et al., 2004; M. Li et al., 2005; W. Li et al., 2003; Matthews et al., 2002; McManus et al., 2006; Pham et al., 2006; Shields et al., 2004). Electrospun protein materials have fiber diameters in the range of those found in native ECM and display improved mechanical properties over hydrogels. The electrospun scaffolds may incorporate additional important ECM components, such as particular subtypes of collagen, glycosaminoglycans, and laminin, either in the spun fibers or as coatings, to promote cell adhesion, growth, and differentiation (Ma et al., 2005; Rho et al., 2006; Zhong et al., 2005). The use of specialized proteins such as silk fibroin offers the opportunities to design scaffolds with enhanced strength or other favorable features (Ayutsede et al., 2006; Jin et al., 2004; Kim et al., 2005; Min et al., 2004), while the use of inexpensive materials such as wheat gluten may enable the production of lower-cost electrospun biomaterials (Woerdeman et al., 2005).
Electrospinning technology also facilitates the production of scaffolds blending proteins with synthetic polymers to confer desired properties. Blending of collagen type I with biodegradable, elastomeric poly(ester urethane)urea generated strong, elastic matrices with improved capacity to promote cell binding and expression of specialized pheno-types as compared to the synthetic polymer alone (W. He et al., 2005; Kwon and Matsuda, 2005; Stankus et al., 2004). Novel properties not normally associated with the ECM may be introduced. For example, nanofibers coelectrospun from polyaniline and gelatin yielded an electrically conductive scaffold with good biocompatibility (M. Li et al., 2006).
One demanding application of scaffold technology is in the production of a biological vascular substitute (Niklason et al., 1999). Electrospun combinations of collagen and elastin or collagen and synthetic polymers have been considered for the development of vascular scaffolds (Boland et al., 2004; W. He et al., 2005; Kwon and Matsuda, 2005; Ma et al., 2005). Recently, electrospinning was utilized to fabricate scaffolds blending collagen type I and elastin with PLGA for use in neo-blood vessels (Stitzel et al., 2006). These scaffolds showed compliance, burst pressure, and mechanics comparable to native vessels and displayed good biocom-patibility both in vitro and after implantation in vivo. When seeded with endothelial and smooth muscle cells, such scaffolds may provide a basis to produce functional vascular grafts suitable for clinical applications such as cardiac bypass procedures.
It may be problematic to introduce cells into a nanofi-brillar structure in which pore spaces are considerably smaller than the diameter of a cell (Lutolf and Hubbell, 2005). However, remarkably, it is possible to utilize electro-spinning to incorporate living cells into a fibrous matrix. A recent proof-of-concept study documented that smooth muscle cells could be concurrently electrospun with an elastomeric poly(ester urethane)urea, leading to "microintegration" of the cells in strong, flexible fibers with mechani cal properties not greatly inferior to those of the synthetic polymer alone (Stankus et al., 2006). The cell population retained high viability, and, when maintained in a perfusion bioreactor, the cellular density in the electrospun fibers doubled over four days in culture. In a similar vein, it has been found that cells can survive inkjet printing (Nakamura et al., 2005; Roth et al., 2004; Xu et al., 2005). Printing of cells together with matrix biomaterials will allow the production of three-dimensional structures that mimic the architectural complexity and cellular distribution of complex tissues. The technology can be applied even to highly specialized, fragile cells, such as neurons. After inkjet printing of hippocampal and cortical neurons, the cells retained their specialized phenotype, as judged by both immunohistochemical staining and whole-cell patch-clamping, a stringent functional test of electrical excitability (Xu et al., 2006). Incorporation of cells by elec-trospinning or printing generates, in a sense, the ultimate smart biomaterials.
At the chemical level, a number of groups have begun to explore the production of biomaterials that unite the advantages of smart synthetic polymers with the biological activities of proteins. The notion of smart polymers initially described materials that show large conformational changes in response to small environmental stimuli, such as temperature, ionic strength, pH, and light (Galaev and Mattiasson, 1999; Williams, 2005). The responses of the polymer may include precipitation or gelation, reversible adsorption on a surface, collapse of a hydrogel or surface graft, and alternation between hydrophilic and hydropho-bic states (A. S. Hoffman et al., 2000). In many cases the change in the state of the polymer is reversible. Biological applications of this technology currently under development span diverse areas, including bioseparation, drug delivery, reusable enzymatic catalysts, molecular switches, biosensors, regulated protein folding, microfluidics, and gene therapy (Roy and Gupta, 2003). In tissue engineering, smart polymers offer promise for revolutionary improvements in scaffolds. Beyond the physical properties of polymers, a major goal is to invest smart biomaterials with specific properties of signaling proteins, such as ECM components and growth factors.
One approach is to link smart polymers to proteins (A. S. Hoffman, 2000; A. S. Hoffman et al., 2000). The proteins can be conjugated either randomly or in a site-specific manner, through engineering of the protein to introduce a reactive amino acid at a particular position. If a conjugation site is introduced near the ligand-binding domain of a protein, induction of a change in conformational state of the smart polymer can serve to regulate the protein's activity (Stayton et al., 1995). This may allow selective capture and recovery of specific cells, delivery of cells to a desired location, and modulation of enzymes, such as matrix metallo-proteases, that influence tissue remodeling.
More broadly, the design of genetically modified proteins or of hybrid polymers incorporating peptides and protein domains will enable the creation of a wealth of novel biomaterials that also can be designated smart (Anderson et al., 2004a). These include engineered mutant variants of existing proteins, semisynthetic scaffold materials incorporating protein domains, scaffold materials linked to synthetic peptides, and engineered peptides capable of self-assembly into nanofibers.
Genetic engineering may improve on natural proteins for applications in tissue engineering (van Hest and Tirrell, 2001). For example, a collagen-like protein was generated by using recombinant DNA technology to introduce tandem repeats of the domain of human collagen II most critically associated with the migration of chondrocytes (Ito et al., 2006). When coated onto a PLGA scaffold and seeded with chondrocytes, the engineered collagen was superior to wildtype collagen II in promoting artificial cartilage formation. Similarly, recombinant technology has been employed to generate a series of elastin-mimetic protein triblock copolymers (Nagapudi et al., 2005). These varied broadly in their mechanical and viscoelastic properties, offering substantial choices for the production of novel materials for tissue engineering.
The incorporation of bioactive signals into scaffold materials of the types just described can be accomplished by the chemical linkage of synthetic peptides as tethered ligands. Numerous studies have confirmed that incorporation of the integrin-binding motif arginine-glycine-aspartic acid (RGD), first identified in fibronectin (Ruoslahti and Pier-schbacher, 1987), enhances the binding of many types of cells to a variety of synthetic scaffolds and surfaces (Alsberg et al., 2002; Hersel et al., 2003; Liu et al., 2004). The CS5 cell-binding domain of fibronectin (Mould et al., 1991) also has been incorporated into scaffolds and its activity shown to be subject to regulation by sequence context (Heilshorn et al., 2005). It is likely that greater selectivity and potency in cellular binding and enhancement of growth and function will be achieved in the future by taking advantage of the growing understanding of the role of additional binding motifs in addition to and/or in concert with RGD (Salsmann et al., 2006; Takagi, 2004). The integrin family comprises two dozen heterodimeric proteins, so there is great opportunity to expand the set of peptide-binding motifs that could be utilized on tissue-engineering scaffolds, with the hope of achieving greater selectivity and control.
The modification of matrices with bioactive peptides and proteins can extend well beyond binding motifs to promote cell adhesion (Boontheekul and Mooney, 2003). Cells also need to migrate in order to form remodeled tissues. Thus, the rate of degradation of scaffolds used for tissue engineering is a crucial parameter affecting successful regeneration (Alsberg et al., 2003). Regulation of the degradation rate can be achieved by varying physical parameters of the scaffold. Alternatively, target sites for proteolytic degradation can be built into the scaffold (Halstenberg et al., 2002; S. H. Lee et al., 2005; Mann et al., 2001). For example, the incorporation into a cross-linked synthetic hydrogel of target sequences for matrix metalloproteases known to play an important role in cell invasion was shown to enhance the migration of fibroblasts in vitro and the healing of bony defects in vivo (Lutolf et al., 2003). Biodegradation of the synthetic matrix was efficiently coupled to tissue regeneration.
Growth factors that drive cell growth and differentiation can be added to the matrix in the form of recombinant proteins or, alternatively, expressed by regenerative cells via gene therapy. Factors of potential importance in tissue engineering and methods to deliver them have been reviewed recently (Vasita and Katti, 2006). Ideally, for optimized tissue formation without risk of hyperplasia, the growth factors should be presented to cells for a limited period of time and in the correct local environment. Biodegradable electrospun scaffolds are capable of releasing growth factors at low rates over periods of weeks to months (Chew et al., 2005; W. He et al., 2005; C. Li et al., 2006). Biologically regulated release of growth factors from scaffolds appears particularly promising as a means to ensure that cells in regenerating neotissues receive these signals when and in the amounts required. For example, by physically entrapping recombinant bone mor-phogenetic protein-2 (BMP-2) in a hydrogel so that it would be released by matrix metalloproteases, Lutolf et al. (2003) achieved excellent bone healing in a critical-size rat calvarial defect model. Similarly, incorporation of a neurotrophic factor in a degradable hydrogel was shown to promote local extension of neurites from explanted retina, and gels were designed to release multiple neurotrophin family members at different rates (Burdick et al., 2006).
Controlled presentation of angiogenic factors such as vascular endothelial growth factor (VEGF) should promote the well-regulated neovascularization of engineered regenerating tissue (Lei et al., 2004; Nomi et al., 2002). Again, it is possible to covalently couple an angiogenic factor to a matrix (Zisch et al., 2001) and to regulate its release based on cellular activity and demand (Zisch et al., 2003). The selection of a sulfated tetrapeptide that mimics the VEGF-binding capability of heparin, a sulfated glycosaminogly-can, provides another potential tool for the construction of scaffolds able to deliver an angiogenic factor to cells in a regulated manner (Maynard and Hubbell, 2005).
Spatial gradients can be generated in the presentation of growth factors within scaffold constructs. This may help to guide the formation of complex tissues and, in particular, to direct migration of cells within developing neotissues (Campbell et al., 2005; DeLong et al., 2005). The introduction of more sophisticated manufacturing technologies, such as solid free-form fabrication, will allow the production of tissue-engineering constructs comprising scaffolds, incorporated cells, and growth factors in precise, complex three-dimensional structures (Hutmacher et al., 2004).
A next stage of smart biomaterials development extends to the design or discovery of bioactive materials not necessarily based directly on naturally occurring carbohydrate or protein structures. At one level this may entail the relatively straightforward chemical synthesis of new materials, coupled with a search for novel activities. By adapting the combinatorial library approach already well established for synthetic peptides and druglike structures, together with even moderately high-throughput assays, thousands of candidate scaffold materials can be generated and tested. Thus, screening of a combinatorial library derived from commercially available monomers in the acrylate family revealed novel synthetic polymers that influenced the attachment, growth, and differentiation of human embryonic stem cells in unexpected ways (Anderson et al., 2004b).
Potentially more revolutionary developments in biomaterials will continue to arise at the interface of tissue engineering with nanotechnology. Basic understanding of the three-dimensional structure of existing biological molecules is being applied to a "bottom-up" approach to generate new, self-assembling supramolecular architectures (Zhang, 2003; Zhao and Zhang, 2004). In particular, self-assembling peptides offer promise because of the large variety of sequences that can be made easily by automated chemical synthesis, the potential for bioactivity, the ability to form nanofibers, and responsiveness to environmental cues (Fairman and Akerfeldt, 2005). Recent advances include the design of short peptides (e.g., heptamers) based on coiled-coil motifs that reversibly assemble into nanofila-ments and nanoropes, without excessive aggregation (Wagner et al., 2005). These smart peptide amphiphiles can be induced to self-assemble by changes in concentration, pH, or level of divalent cations (Hartgerink et al., 2001, 2002). Branched structures can be designed to present bioactive sequences such as RGD to cells via nanofiber gels or as coatings on conventional tissue-engineering scaffolds (Guler et al., 2006; Harrington et al., 2006). In addition, assembly can occur under conditions that permit the entrapment of viable cells in the resulting nanofiber matrix (Beniash et al., 2005). The entrapped cells retain motility and the ability to proliferate.
Further opportunities exist to expand the range of peptidic biomaterials by utilizing additional chemical components, such as porphyrins, which can bind to pep-tides and induce folding (Kovaric et al., 2006). Porphyrins and similar structures also may add functionality, such as oxygen storage, catalysis or photosensitization of chemical reactions, or transfer of charge or molecular excitation energy.
Peptide-based nanofibers may be designed to present bioactive sequences to cells at very high density, substantially exceeding that of corresponding peptide epitopes in biological ECM. For example, a pentapeptide epitope of laminin, isoleucine-lysine-valine-alanine-valine (IKVAV), known to promote neurite extension from neurons, was incorporated into peptide amphiphiles (PA) capable of self-assembly into nanofibers that form highly hydrated (>99.5 weight % water) gels (G. A. Silva et al., 2004). When neural progenitor cells capable of differentiating into neurons or glia were encapsulated during assembly of the nanofibers, they survived over several weeks in culture. Moreover, even without the addition of neurotrophic growth factors, they displayed neuronal differentiation, as exemplified by the extension of large neurites, already obvious after one day, and by expression of piII-tubulin. The production of neuronlike cells from the neural progenitors, whether dissociated or grown as clustered "neurospheres," was more rapid and robust in the IKVAV-PA gels than on laminin-coated substrates or with soluble IKVAV. By contrast, the production of cells expressing glial fibrillary acidic protein (GFAP), a marker of astrocytic differentiation, was suppressed significantly in the IKVAV-PA gels, even when compared to growth on laminin, which favors neuronal differentiation. The ability to direct stem or progenitor cell differentiation via a chemically synthesized biomaterial, without the need to incorporate growth factors, offers many potential advantages in regenerative medicine.
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