The main naturally derived biomaterials used as scaffolding for tissue engineering are of varying chemical nature. They comprise polypeptides, polysaccharides, polyesters, and inorganic materials. A mammalian body has different kinds of polypeptides including plasma, structural, and functional proteins. The majority of proteins used as biomaterials originate from blood plasmas and structural skeletons. Functional proteins such as enzymes, cell growth factors, and interleukins are also used, but are mostly incorporated into biomaterials as ingredients. Here we have focused our attention on the most frequently used natural scaffolds for tissue engineering, such as fibronectin and collagen.
Fibronectin is a multifunctional component of the ECM. Intracellular signaling induced by cell adhesion on fibronectin plays a critical role in cytoskeletal organization, cell cycle progression, and cell survival (Hynes, 1990; Frisch and Ruoslahti, 1997). Cells assemble fibronectin into a fibrillar form that accumulates at their apical surface. Fibronectin matrix formation is initiated by fibronectin binding to cell surface receptors, followed by assembly and reorganization of the cell surface–associated fibronectin into fibrils (McDonald, 1988; Mosher et al., 1992; Mosher, 1993). The α5β1 integrin is the major receptor responsible for fibronectin matrix assembly (Ruoslahti, 1991; Wu et al., 1993). The αvβ3 integrin can also direct matrix assembly (Wennerberg et al., 1996; Wu et al., 1996), which may account for fibronectin matrix formation in α5 integrin null mice (Yang et al., 1993). A number of other integrins bind to fibronectin, but do not initiate fibril formation (Busk et al., 1992; Zhang et al., 1993; Wu et al., 1995). The first type III repeat of the fibronectin molecule is important in promoting matrix assembly (Morla and Ruoslahti, 1992; Aguirre et al., 1994; Hocking et al., 1994). Small fragments derived from this III1 module induce fibronectin polymerization at moderate concentrations, but inhibit it at high concentrations (Morla and Ruoslahti, 1992; Morla et al., 1994). Interaction of integrins with the actin cytoskeleton is also essential for matrix assembly (Hynes, 1990).
Bourdoulous et al. studied the role of the fibronectin matrix in cytoskeletal regulation by disassembling the fibronectin matrix with a 76-amino acid fragment derived from the first type III repeat of fibronectin (Morla and Ruoslahti, 1992; Morla et al., 1994), or by preventing matrix assembly with integrin antibodies. Interestingly, they found that removal of fibronectin matrix without altering the cell-substrate adhesion of the cells decreased the basal activity of ERK while slightly potentiating ERK response to growth factors. This effect and the activation of P38 MAPK and suppression of JNK in III1-C-treated cells showed that the effects of matrix removal on MAPKs are quite different from those of loss of substrate adhesion. In agreement with this, depleting fibronectin matrix had little effect on cell survival, while inhibiting cell proliferation. Thus, signals from fibronectin matrix seem to control cell proliferation, whereas cell substrate adhesion provides a survival signal. These signals can be modulated separately by removing the matrix or by allowing cells to attach to a substrate in the absence of matrix.
In summary, these results revealed a specialized role for cell surface fibronectin matrix in cytoskeletal organization, growth factor responses, and cell cycle control that cannot be substituted for by cell adhesion to a substrate.
Furthermore, it is known that α5β1 and αvβ3 integrins are central to regulating downstream events, including cell survival and cell-cycle progression. In contrast to previous findings that αvβ3 integrins promote angiogenesis (Bourdoulous et al., 1998), recent evidence argues that αvβ3 integrins may act as negative regulators of proangiogenic integrins such as α5β1. This suggests that fibronectin is critical for scaffold vascularization because it is the only mammalian adhesion protein that binds and activates α5β1 integrins. Cells are furthermore capable of stretching fibronectin matrices such that the protein partially unfolds, and recent computational simulations provide structural models of how mechanical stretching affects fibronectin function. Vogel and Baneyx (2003) recently proposed a model, whereby excessive tension generated by cells in contact with biomaterials may in fact render fibronectin fibrils non-angiogenic and may potentially inhibit vascularization. The model could explain why current biomaterials fail to vascularize, independent of their surface chemistries and textures.
In addition, other factors seem to modulate fibronectin fibril structure and thereby affect angiogenesis. Hall et al. (2001) showed that the molecular properties of fibrin-based matrices, such as fibrillar structure and covalent modifications with adhesion domains, influence the angiogenic behavior of human umbilical vein endothelial cells (HUVECs) in vitro. The fibrillar structure of fibrin-based matrices was influenced by pH but not by covalent incorporation of exogenous adhesion domains. Native fibrin-based matrices polymerized at pH 10 formed organized and longitudinally oriented fibrin fibrils, which provided a good angiogenic substrate for endothelial cells. Furthermore, upon covalent incorporation of the model ligand L1Ig6, which binds to the integrin most prominently expressed on the surface of angiogenic endothelial cells, alpha(v)beta3, these matrices became angiogenesis-promoting when polymerized at physiological pH. Most important, L1Ig6-modified matrices were very specific in inducing the angiogenic phenotype of HUVECs, whereas control cells did not differentiate on these matrices. These results indicate that artificial ECMs can influence cell behavior in two ways. One way is based on the three-dimensional fibril structure of the matrix molecules themselves, and the other provides specific binding sites for direct cell-matrix interactions that lead to the activation of second-messenger cascades, thus promoting angiogenic differentiation.
Moreover, starting from the observation that currently used biodegradable scaffolds in cardiovascular tissue engineering show toxic degradation and inflammatory reactions and are potentially immunogenic, Ye et al. (2000) proposed the use of a three-dimensional fibrin gel scaffold for vessel tissue engineering. In their experiments, human aortic tissue was harvested from the ascending aorta in the operation room and worked up to pure human myofibroblast cultures. These human myofibroblast cultures were suspended in fibrinogen solution and seeded into 6-well culture plates for cell development for 4 weeks and supplemented with different concentrations of aprotinin. Hydroxyproline assay and histological studies were performed to evaluate the tissue development in these fibrin gel structures. The light microscopy and the transmission electron microscopy studies for tissue development based on the three-dimensional fibrin gel structures showed homogenous cell growth and confluent collagen production. No toxic degradation or inflammatory reactions could be detected. Furthermore, fibrin gel myofibroblast structures dissolved within 2 days in medium without aprotinin, but medium supplemented with higher concentration of aprotinin retained the three-dimensional structure and had a higher collagen content and a better tissue development. They concluded that a three-dimensional fibrin gel structure could serve as a useful scaffold for tissue engineering with controlled degradation, excellent seeding effects, and good tissue development.
Collagens are ubiquitous proteins responsible for maintaining the structural integrity of vertebrates and many other organisms (Myllyharju and Kivirikko, 2001). More than 20 genetically distinct collagens have been identified (Hulmes, 1992, 2002; Kadler et al., 1996; Ottani et al., 2001). In tissues that have to resist shear, tensile, or pressure forces, such as tendons, bone, cartilage, and skin, collagen is arranged in fibrils, with a characteristic 67 nm axial periodicity, which provides the tensile strength. Only collagen types I, II, III, V, and XI self-assemble into fibrils. The fibrils are composed of collagen molecules, which consist of a triple helix of approximately 300 nm in length and 1.5 nm in diameter. Collagen fibril formation is an extracellular process, which occurs through the cleavage of terminal procollagen peptides by specific procollagen metalloproteinases.
Some collagens form networks (types IV, VIII, and X), a typical example of which is the basement membrane, mostly made of collagen IV. Other collagens associate with fibril surfaces (types VI, IX, XII, and XIV). Yet other collagens are transmembranous proteins (types XIII and XVIII) or form periodic beaded structures (type VI).
Type I collagen occurs throughout the body, except in cartilage. It is the principal collagen in the dermis, fasciae, and tendons and is a major component of mature scar tissue. Type II collagen occurs in cartilage, the developing cornea, and in the vitreous body of the eye. Type III collagen dominates in the wall of blood vessels and hollow intestinal organs and co-polymerizes with type I collagen. Types V and XI collagen are minor components and occur predominantly co-polymerized with collagen I (type V) and collagen II (type XI).
Collagens are mostly synthesized by the cells comprising the ECM: fibroblasts, myofibroblasts, osteoblasts, and chondrocytes. Some collagens are also synthesized by adjacent parenchymal or covering (epithelial, endothelial, and mesothelial) cells. A typical example is type IV collagen, which is synthesized in a cooperative effort between the stromal cell and the parenchymal/covering cell.
Collagen is the most abundant protein in animals and because of its high mechanical strength and good resistance to degradation, it has been utilized in a wide range of products in industry (Ulrich et al., 1992), while its low antigenicity has resulted in its widespread use in medicine (Ramshaw et al., 2000). Collagen products can be purified from fibers, from molecules reconstituted as fibers, or from specific recombinant polypeptides with preferred properties. A feature common to all these biomaterials is the need for stable chemical cross-linking to control the mechanical properties and the residence time in the body, and to some extent the immunogenicity of the device. This can be achieved by a number of different cross-linking agents that react with specific amino acid residues on the collagen molecule imparting individual biochemical, thermal, and mechanical characteristics to the biomaterial (Silver et al., 1995).
For these reasons, collagen is currently used for tissue engineering. For example, Hudon et al. (2003) produced a new model of endothelialized reconstructed dermis that promotes the spontaneous formation of a human capillary-like network. The endothelialized dermis was prepared by co-culturing two human cell types, dermal fibroblasts and umbilical vein endothelial cells, in a collagen sponge biomaterial. Thereafter, they strove to study, quantitatively and qualitatively, the influence of angiogenic and angiostatic drugs on capillary-like tube (CLT) formation in vitro in the model. The visualization by confocal microscopy of the tubes present in the model showed that the endothelial structures were not cord-like but rather CLTs with well-defined lumina. Moreover, these tubes were organized in a complex network of branching structures. When angiogenic factors (vascular endothelial growth factor (VEGF) 10 ng/ml or basic fibroblast growth factor 10 ng/ml) were added to the model, 1.8 and 1.4 times more capillaries were observed, respectively, whereas the addition of progesterone (10 μg/ml) reduced by 2.4 times the number of tubes compared with the control. These results suggest that this model is a highly efficient assay for the screening of potentially angiogenic and angiostatic compounds.
Another example of collagen as scaffold for dermal tissue engineering came from recent studies by Guerret et al. (2003). They studied Apligraf, a bioengineered living skin, composed of a bovine collagen lattice containing living human fibroblasts overlaid with a fully differentiated epithelium made of human keratinocytes. To investigate its progressive remodeling, athymic mice were grafted and the cellular and the ECM components were studied from 0 to 365 days after grafting. Biopsies were analyzed using immunohistochemistry with species-specific antibodies and electron microscopy techniques. They observed that this bioengineered tissue provided living and bioactive cells to the wound site up to 1 year after grafting. The graft was rapidly incorporated within the host tissue and the bovine collagen present in the graft was progressively replaced by human and mouse collagens. A normal healing process was observed, that is, type III collagen appeared transiently with type I collagen, the major collagen isoform present at later stages. New molecules, such as elastin, were produced by the living human cells contained within the graft. This animal model combined with species-specific immunohistochemistry tools is, thus, very useful for studying long-term tissue remodeling of bioengineered living tissues.
Jux et al. (2003) used intestinal collagen layer (ICL), a highly purified (acellular) bioengineered type-I collagen derived from porcine submucosa, as septal occluder to replace a commercial occluder (CardioSeal) in percutaneous transcatheter closure of interventionally created atrial septal defects in lambs. A complete pathomorphological follow-up investigation including histology was carried out after 2, 4, and 12 weeks. Standard CardioSEAL implants served as a control group. After 2 weeks in vivo the devices were already covered completely by neo-endothelium. Compared with the conventional synthetic scaffold, ICL devices showed a quicker endothelialization, decreased thrombogenicity, and superior biocompatibility with no significant cellular infiltration observed in the histology of explants with ICL fabrics. After 3 months in vivo the collagen layer remained mechanically intact, but began to show the first histological signs of mild disintegration, gradual reabsorption, and remodeling. In conclusion, short-term results from preliminary in vivo experiments using a bioengineered collagen matrix as the occluder tissue scaffold showed excellent biocompatibility. This resulted in superior overall results: quicker endothelialization, decreased thrombogenicity, and decreased immunological host response. Another recent clinical application of collagen was provided by Sculean et al. (2003). They compared, clinically, the treatment of deep intrabony defects with a combination of a bovine-derived xenograft (BDX) and a bioresorbable collagen membrane to access flap surgery.
Twenty-eight patients suffering from chronic periodontitis, each of whom displayed one intrabony defect, were randomly treated with BDX + collagen membrane (test) or with access flap surgery (control). Soft tissue measurements were made at baseline and at 1 year following therapy. At 1 year after therapy, the test group showed a reduction in mean probing depth (PD) from 9.2 ± 1.3 to 3.9 ± 0.7 mm and a change in mean clinical attachment level (CAL) from 10.2 ± 1.5 to 6.2 ± 0.5 mm. In the control group, the mean PD was reduced from 9.0 ± 1.2 to 5.2 ± 1.8 mm and the mean CAL changed from 10.5 ± 1.5 to 8.4 ± 2.1 mm. The test treatment resulted in statistically higher PD reductions and CAL gains than the control one. In the test group all sites (100%) gained at least 3 mm of CAL. In the control group no CAL gain occurred in four sites (29%), whereas at six sites (43%) the CAL gain was 2 mm. A CAL gain of 3 mm or more was measured in four defects (29%). From these studies, it can be concluded that at 1 year after surgery both therapies resulted in significant PD reductions and CAL gains, and treatment with BDX + collagen membrane resulted in significantly higher CAL gains than treatment with access flap surgery.
While these examples offer encouraging applications of collagen in tissue engineering, limits in its use arise from its low mechanical strength and a fast biodegradation when implanted in the human body. Thus, more research is needed to improve mechanical strength and to enhance the time of permanence in the human body.