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Alginate is a polysaccharide derived from brown seaweed, which has the unique property of being able to form a gel in the presence of certain divalent cations (e.g., calcium, strontium, or barium). Alginate has been of commercial interest in the food industry since the 1930s, when its properties as an emulsifier, thickener, and stabilizer were recognized. Alginate has also long been used for biomedical purposes, particularly in the manufacture of surgical dressings for exuding wounds (1). However, the explosive increase in medical applications for alginate, began with the recognition of its use as a scaffold for the encapsulation and immunoprotection of transplanted cells. The encapsulation of non-autologous islet cells in alginate for the treatment of diabetes is based on the concept that nutrients can diffuse in and insulin out of the alginate construct without triggering an immune response (2-3). Similarly, alginate has been used to immunoprotect recombinant cells delivering tumor-suppressing agents (4-5) and growth hormone (6). Stable cultures in alginate beads have been achieved with a number of cell types including chondrocytes, bone-marrow stromal cells, islets, myoblasts, fibroblasts, Schwann cells, kidney cells, epithelial cells, and hepatocytes. For orthopedic purposes, encapsulated bone-marrow stromal cells and chondrocytes have been proposed for the healing bone and cartilage defects (7-9). As a bulking agent, alginate has gained attention as a space-filling material for treating pediatric urinary reflex (10-12). [ABSTRACT FROM AUTHOR]

The exponential growth in basic research and clinical trials involving tissue-engineered materials has generated a corresponding need for the evaluation of the material properties and functional performance of these constructs during development and/or after implantation. Applications focusing on musculoskeletal tissues, in particular, require detailed assessment of the biomechanical properties of neo-tissue constructs in vitro and in vivo (1). Based on the known properties of normal tissues, investigators have identified a range of biological, biochemical, and biophysical end-point parameters that must be quantified to determine the potential for success of a particular tissue-engineering methodology. Such end-point assessment is critical to our understanding of the basic scientific approaches underlying tissue engineering. In addition, biomechanical assessment is crucial for the implementation of regulatory processes that are coupled to clinical practice. [ABSTRACT FROM AUTHOR]

The ultimate goal of tissue engineering is to replace, repair or enhance the biological function of damaged, absent or dysfunctional elements of a tissue or an organ. Engineered tissues are produced by using cells that are manipulated through their extracellular environment to develop living biological substitutes for tissues that are lacking or malfunctioning (1-5). Many different strategies may be used to accomplish this goal. Among the most important factors that determine the selection of the best strategy for developing and utilizing engineered tissues are the technical feasibility, the required properties of the implant, and the interaction of the host with the graft. [ABSTRACT FROM AUTHOR]

Analysis of gene expression at the mRNA level may be used, together with other structural and functional evaluations, to characterize the quality of engineered tissues. In turn, engineered tissues may be used as model systems of developing tissues to investigate how mRNA gene expression is modulated by a variety of factors, including structural (e.g., type of three-dimensional [3D] scaffold), biochemical (e.g., combination of bioactive molecules), and physical (e.g., regime of bioreactor cultivation). In order to reliably evaluate mRNA levels, Northern blotting and in situ hybridization (ISH) methods are not sensitive enough to detect low-level gene expression, and not accurate enough to quantify the full range of expression. An amplification step is thus often required to quantify mRNA amounts from engineered tissues of limited size. Quantification of mRNA using conventional reverse-transcriptase polymerase chain reaction (RT-PCR) techniques may be of limited accuracy, both because of the detection method used (e.g., semi-quantitative image analysis), and because analysis is often performed using a constant number of amplification cycles, after which the system could be in a saturation phase. The recently established real-time RT-PCR technology has made mRNA quantitation more reproducible, precise, and sensitive than conventional RT-PCR, because it allows measurement of the amount of amplified product using a quantitative laser-based method and in the early exponential phase of the PCR reaction, when none of the reagents is rate-limiting (1). [ABSTRACT FROM AUTHOR]

There have been significant advances in cartilage-engineering methodology in recent years, particularly in the production of biodegradable scaffolds and the use of bioreactors to improve the outcome of the engineering process. To date, engineered cartilage studies have mostly relied on histochemical and immunohistochemical methods for the characterization of the engineered cartilage matrix. The detection of various matrix components using these methods may not be sufficient to judge the quality of engineered cartilage and compare it objectively to natural cartilage. In order for cartilage tissue engineering to progress to the clinic, quantitative methods for the analysis of different matrix components will be required to monitor and ensure the quality of engineered cartilage. [ABSTRACT FROM AUTHOR]

Visualizing and quantifying cellular contact with materials is a critical step in the evaluation of cell-material interaction both in vitro and in vivo. To achieve this, a number of techniques are available. The focus of this chapter is the presentation of staining techniques for use with light microscopes, including transmitted, reflected, fluorescence, or laser-scanning confocal. Other visualization techniques and the potential of techniques such as transmission electron microscopes (TEM), scanning electron microscopy (SEM), and atomic force microscopy (AFM) should not be ignored. But the subject becomes too broad and complex to enable a single protocol to be all-encompassing. However, many of the protocols described can be modified easily for use with TEM and SEM visualization. [ABSTRACT FROM AUTHOR]

Numerous publications have described the histology of tissue-engineered constructs and tissue-biomaterial interfaces observed with the light microscope (1-5). Indeed, this approach has become a routine method for the evaluation of the biological quality of engineered tissues. Additional microscopical techniques, including immunohistochemistry and confocal microscopy, have also been described. These methods have enabled the localization of important biological molecules in the tissue-engineered construct. However, all of these techniques are limited to describing relatively gross structural features, and they give little insight into the ultrastructural relationships between cells and biomaterials within a construct. Scanning electron microscopy (SEM) has been applied more recently to the study of constructs, but it too provides only limited information and not the "fine detail" of cell-biomaterial interaction (6,7). Despite the limitations noted here, relatively little research has been reported in which the transmission electron microscope (TEM) has been used to study the ultrastructure of cell-polymer constructs. This may be partly because of the difficulties associated with specimen preparation when living tissues and materials are both present. However, these difficulties are not insurmountable. Studies of the cell-biomaterial interface using electron microscopy have been reported in the general field of biomaterials and medical devices (8-11). Here, they have enabled researchers to gain a greater understanding of the response of cells and tissues to biomaterial surfaces. [ABSTRACT FROM AUTHOR]

Examination of engineered tissues is clearly necessary to evaluate their properties and determine how close they are to those of the original tissue being engineered. Microscopy encompasses a group of techniques that allows the assessment of many parameters, including information on the cells themselves, on their viability, density, proliferation status, morphology, their capacity for protein synthesis, and their cell activity. Microscopy can also be utilized to follow the structure of any scaffold used and to determine whether it is broken down. In addition to examining how the cells interact with it, the scaffold can be examined to establish whether they adhere successfully, which molecules are involved, and whether the cells go on to produce any matrix—and if so, whether it replaces the original scaffold or integrates with it and builds on it. [ABSTRACT FROM AUTHOR]

Tissue engineering is a major focus of biotechnological research today, with the expectation that this type of technique will ultimately transform medical practice (1). The most ambitious tissue-engineering schemes assume that specific tissues and organs will be restored, in a multistage fabrication procedure. For example, cells derived from the patient may be processed to increase the total number available, seeded into a suitable three-dimensional (3D) resorbable scaffold and further processed in vitro to induce the elaboration of neo-tissue prior to implantation (2). [ABSTRACT FROM AUTHOR]

The chondrocyte is responsible for maintaining the integrity of the complex extracellular matrix (ECM) of articular cartilage. Therefore, chondrocytes directly control the biomechanical properties of the tissue that enable articular cartilage to bear cyclical compressive loading (1). The ultimate goal of articular cartilage engineering is to produce a cartilage construct with identical structure and properties to native tissue. Although this goal has not yet been achieved, many advances have been made in understanding the biology of the extracellular cartilage matrix (1-5). [ABSTRACT FROM AUTHOR]

Engineered tissues could form the basis for novel therapies for millions of patients who suffer from the loss of tissue or its function (1), and be used for in vitro studies of tissue development and normal and pathological function. One approach to tissue engineering involves seeding of a high density of uniformly distributed cells on three-dimensional (3D) polymeric scaffolds, and cultivating the resulting cell-polymer constructs under conditions that permit the formation of functional tissues (2). The scaffold provides a defined structure for cell attachment and tissue development, and the bioreactor provides control over the biochemical and physical factors in the cell environment. [ABSTRACT FROM AUTHOR]

One of the most widely studied bone-tissue engineering approaches involves the seeding and extended in vitro culturing of cells within a biodegradable scaffold prior to implantation. The bioresorbable scaffold must be biocompatible and porous in order to facilitate rapid vascularization and growth of newly formed tissue (1-8). During the in vitro culture period, the seeded cells proliferate and secrete tissue-specific extracellular matrix (ECM). The selection of the scaffold material, which should exhibit an adequate three-dimensional (3D) porous structure, is a primary consideration in matrix- and cell-based bone-tissue engineering strategies (9). [ABSTRACT FROM AUTHOR]

Bioactive polymeric materials can be constructed via the modification of bioinert polymers with bioactive peptides such as cell-adhesion peptides or growth factors. Such materials can be designed to induce specific biological responses desired in applications such as tissue engineering and regenerative medicine. For example, transforming growth factor-beta (TGF-β) has been immobilized to polymeric scaffold materials to increase the synthesis of extracellular matrix (ECM) proteins and thus improve the mechanical properties of the engineered tissue (1), and epidermal growth factor has been immobilized to improve the function of hepatocytes (2). Cell adhesion peptides have been grafted to polymeric scaffold materials in order to enhance and promote cell attachment and spreading (3-5). Alternatively, materials can be designed to be cell-selective, and adhesive for only a particular cell type, by modifying cell non-adhesive polymers with cell-selective adhesion ligands. Examples of cell-selective ligands include the peptide REDV, which interacts with endothelial cells but not platelets, fibroblasts, or smooth-muscle cells (SMCs) (6), and the peptide KRSR, which interacts with osteoblasts but not fibroblasts (7). [ABSTRACT FROM AUTHOR]

Current research in the field of tissue engineering is focused on the development of appropriate strategies for repair and regeneration of biological tissues. Biological tissues consist of cells situated within a complex molecular framework known as the extracellular matrix (ECM) with an integrated vascular system for oxygen or nutrient supply. In soft tissues, native ECMs consist mainly of collagens, proteoglycans, glycosaminoglycans (GAGs), laminins, and fibronection (1). [ABSTRACT FROM AUTHOR]

Bone morphogenetic proteins (BMPs) are members of the transforming growth factor-β (TGF-β) superfamily of proteins that have crucial roles in growth and regeneration of skeletal tissues. In particular, recombinant human bone morphogenetic protein-2 (rhBMP-2) has been shown to induce new bone formation by differentiation of mesenchymal progenitor cells into osteoblasts (1). BMPs act locally, and thus for a clinically beneficial outcome, a carrier matrix is required to retain therapeutic levels of BMP at the repair site (2,3). The potent osteoinductive ability of BMP-2 has been demonstrated in numerous preclinical (4-9) and clinical studies (10-14). Recent data from clinical trials have reported success in the areas of open fracture repair (10,14), interbody spinal fusion (11), and maxillofacial reconstruction (12,13) using rhBMP-2 delivered in an absorbable collagen sponge (ACS). Because rhBMP-2/ACS requires implantation, there is a continuing need for an easily injectable carrier to deliver rhBMP-2 percutaneously to closed fractures. Here, we outline characterization of a calcium phosphate (CaP)-based injectable carrier as a delivery matrix for rhBMP-2. [ABSTRACT FROM AUTHOR]

Chitin, poly(N-acetylglucosamine) is a structural polysaccharide found in the shell of crustaceans and mollusks and is typically isolated from waste produced in food processing. The procedure for extracting chitin from crustacean shells usually involves decalcification, deproteination, and decolorization, which may be accomplished by sequential treatment in 5% HCl solution, 5% NaOH, and 0.5%KMnO4(aq) and oxalic acid. Chitosan, poly(N-acetyl-2-amino-2-deoxy-d-Gluco-pyranose), is formed through the N-deacetylation of chitin, usually by treatment with strong alkalis (e.g., 50% NaOH) at an elevated temperature. It is a linear, polycationic copolymer of N-acetylglucosamine and glucosamine (combined by 1-4 glycosidic linkages), with the degree of deacetylation defined as the molar fraction of N-acetylglucosamine units. Several studies have linked the biological activity of chitosan to its degree of deacetylation and mol wt (1-6), parameters that may be controlled by varying the time and temperature of the deacetylation reaction. [ABSTRACT FROM AUTHOR]

Hyaluronan (HA) is a naturally occurring, negatively charged high mol-wt glycosaminoglycan (GAG), that is composed of repeated disaccharide units of D-glucuronic acid and N-acetylglucosamine. It is the only GAG that lacks an associated protein moiety and sulfate groups. HA is a highly conserved and widely distributed polysaccharide. In a variety of mammalian tissues, it exerts structural functions because of its peculiar physicochemical properties. Because of its propensity to form highly hydrated and viscous matrices, HA imparts stiffness, resilience, and lubrication to tissues. The unique properties of HA are manifested in its mechanical function in the synovial fluid, the vitreous humor of the eye, and the ability of connective tissue to resist compressive forces, as in articular cartilage. [ABSTRACT FROM AUTHOR]

Fibrinogen exerts adhesive effects on cultured fibroblasts and other cells. Specifically, fibrin(ogen) and its various lytic fragments (e.g., FPA, FPB, fragments D and E) were shown to be chemotactic to macrophages, human fibroblasts, and endothelial cells (1-3). Thrombin has also been shown to exert proliferative and adhesive effects on cultured cells (4-7). We previously demonstrated that covalently coating inert Sepharose beads with either fibrinogen or thrombin rendered them adhesive to a wide range of cell types. We employed such coated Sepharose beads to screen or rank normal and transformed cells for their haptotactic responses to fibrinogen (8,9). [ABSTRACT FROM AUTHOR]

Poly (α-hydroxyacids) were found to be bioabsorble and biocompatible in the 1960s (1,2). They are the most widely known, studied and used bioabsorbable synthetic polymers in medicine. Polyglycolide (PGA) and poly-l-lactide (PLLA) homopolymers and their copolymers (PLGA), as well as polylactic acid stereocopolymers produced using l-, d-, or DL-lactides and rasemic polymer copolymer PLDLA are all poly (α-hydroxyacids) (3). Poly (α-hydroxy acids) can be polymerized via condensation, although only low mol-wt polymers are produced. In order to obtain a higher mol wt and thus mechanical strength and longer absorption time, the polymers are polymerized from the cyclic dimers via ring-opening polymerization using appropriate initiators and co-initiators. The most commonly used initiator is stannous octoate (2,3). [ABSTRACT FROM AUTHOR]