Biodegradable synthetic polymers
Easy to process
Less biological properties than natural polymers
Controlled biodegradation rate
Degrade to form by-products
Controlled surface characteristics
Tunable mechanical properties
Difficult to control biodegradation rate
Mimic natural ECM structure and composition
Poor mechanical stability
Transfer of pathogens possible
Acellular tissue matrices
Appropriate biomechanical properties to apply
Similar structure and composition in native tissues
Difficult to re-cellularize
Transfer of pathogens possible
Mostly natural polymers and synthetic polymers
Limited fabrication processing
Synergic effects can be expected
In some cases, poor interaction of interface
Composition can be controlled
2.2.1 Natural Polymers
Natural polymers are defined as those that are produced by the cells of a living organism. Particularly, structural proteins such as collagen, laminin, elastin, and fibronectin have been used as scaffold materials for tissue engineering applications and as vehicles for cell delivery. Among these, collagen has been found a common use as a scaffold and carrier for cells in various tissue engineering applications [13–15], because it is the most abundant protein in mammalian tissues as the major component of the ECM materials. Collagen, which has a triple helix structure, mainly provides the structural integrity of vertebrates and organisms. Collagen can be extracted and purified from these animal tissues by enzymatic digestion and salt/acid extraction methods. The sources of collagens are mostly animal tissues such as bovine skin and tendon, porcine skin and tendon, and rat tail. Since the use of animal-derived collagens has been concerned due to the possible transmission of infectious agents, the recombinant human collagens have been produced . Although collagen has excellent biological functions [17, 18], its low biomechanical properties and fast degradation by enzymes as well as antigenicity potential have been limited for clinical applications.
Silk fibroin is an insoluble fibrous protein obtained by silk worms (Bombyx mori), which contains approximately 90% amino acids, including glycine, alanine, and serine [19–21]. Purified silk fibroin can be extracted from degummed silk (boiling off), which refers to the removal of serine. During the past decades, silk fibroin has been commonly used as a suture material (Surusil®, Suru; Sofsilk™, Covidien). To date, various studies reported the use of silk materials in nanofibers, films, scaffolds, hydrogels, and microspheres for tissue engineering applications with various cell types and drug delivery applications [21–23]. The unique physicochemical properties of silk include a hydrophobic structure, strong intramolecular and intermolecular interactions, and crystal polymorphism. Therefore, silk fibroin enables tunable and slow biodegradation kinetics without toxic chemical cross-linkers. The hydrolysis of silk fibroin is mainly attributed to the proteolytic enzymes, including part of the foreign-body response cascade .
Natural carbohydrates have been utilized as hydrogels for tissue engineering applications and drug carrier materials for drug delivery system . For instance, the hydrophilic polysaccharide hyaluronic acid (HA), which is composed of repeating disaccharide units of β-D-glucuronic acid and N-acetyl-β-D-glucosamine, is broadly distributed in the ECM and shows a critical role in vertebrate tissue morphogenesis . Due to its excellent biocompatibility and hydrodynamic features, HA has been approved for use in human patients as viscous fluid and sheet formulations for knee pain control and surgical adhesives, respectively. To date, many clinical trials have confirmed the safety and effectiveness of HA for human clinical trials [27–31]. The bioactivity of HA, like that of other carbohydrates, can be also improved by physical and/or chemical modification to modulate the cell migration, spreading, and multiplication.
Other carbohydrate polymers such as alginate, dextran, and chitosan have been used in various biomedical applications. Alginate (alginic acid), composed of linear block copolymers of [1–4]-linked β-d-mannuronic acid (M) and α-l-guluronic acid (G) covalently linked together in different blocks, has been used extensively in a hydrogel form for cell encapsulation and drug delivery  as well as in tissue engineering applications . Alginate is ionically cross-linked by the addition of divalent cations (e.g., Ca2+) in aqueous solution. The production of purified alginates with consistently high medical-grade quality has been considered. The purity of the input algal source is critical, along with the development of validation strategies for alginate extraction and purification processes. Alginate hydrogel undergoes a slow and unpredictable dissolution process in vivo, mainly due to the sensitivity of the hydrogels toward calcium chelating compounds . Moreover, a key engineering challenge in designing immunoisolation of the alginate-based microcapsules for cell encapsulation is that of maintaining unhampered exchange of nutrients, oxygen, and therapeutic molecules released by the encapsulated cells while avoiding swelling and subsequent rupture of the alginate microcapsules . Although many advances in alginate-based biomaterials have been informed, numerous progresses will be required for clinical applications.
Chitosan is also a linear polysaccharide of [1–4]-linked D-glucosamine and N-acetyl-D-glucosamine residues, which is a polycationic material produced by the deacetylation of chitin. Chitosan is insoluble in aqueous solution (above pH 7) because of its crystalline structure, while it is fully soluble in dilute acids (below pH 5). It has been used in a number of gene and drug delivery applications due to the abundant positive charged functional groups. Also, its applications in tissue engineering and regenerative medicine have been broadly investigated [36–38].
2.2.2 Acellular Tissue Matrices
ECM contributes to biomechanical/structural integrity and signaling and regulatory functions in the development, maintenance, and regeneration of tissues. ECM components participate in the tissue-specific control of gene expression through a variety of transduction mechanisms in synergic effects with soluble biochemical signals provided by growth factors, cytokines, and hormones [39–41]. Moreover, ECM is itself a dynamic structure that is actively remodeled by direct interactions with the cells . Development of novel scaffolding systems that more closely recapitulate the biological and biomechanical functions of native ECM remains an important area of tissue engineering approaches . Understanding ECM’s complex functions in mature or regenerating tissues continues to be an important task to design a synthetic microenvironment.
In the current limitations for de novo construction of a true ECM mimic from purified ECM components, decellularized tissue matrices have been considered as an ideal scaffolding material due to their biological, structural, and biomechanical similarity to native tissues and their possession of tissue-specific ECM proteins which remain after the decellularization process . As a result, these matrices have been used for a number of tissue engineering applications and have yielded an understanding of the physicochemical properties of ECM as well as a tissue-specific scaffold for engineering functional tissues or organs. For example, by harnessing the cell-matrix interactions that are crucial for cellular behaviors, including adhesion, proliferation, and differentiation , numerous studies have encouraged the tissue formation by combining cells with tissue-specific ECM scaffolds from decellularized heart , urinary bladder , liver , lung , kidney , and small intestinal submucosa (SIS) . Each of these decellularized tissue matrices maintained the biomechanical properties and structural integrity of the original tissue or organ with minimal disruption to the ECMs. More importantly, these ECM proteins play a vital role in tissue maintenance and regeneration of a specific tissue type, and they can modulate cell adhesion and migration, growth factor storage and release, and progenitor cell activation and differentiation [51, 52].
One of the most common acellular tissue matrices is small intestinal submucosa (SIS), which is prepared by mechanical removal of tunica mucosa and muscularis, and serosal layer . Acellular SIS contains various bioactive molecules such as fibronectin, glycosaminoglycans (GAGs), and growth factors . Recently, commercially available SIS-based matrices have been successfully implanted in several reconstructive surgical procedures, which include hernia repair [55, 56], abdominal wall closure , urinary incontinence [58, 59], bladder augmentation , pelvic organ prolapse repair , and so on. As SIS provides high biological variability, standardized protocol for preparation, and manufacturing procedure with quality control, SIS-based tissue products can be a reliable therapeutic option for clinical applications.
Skeletal muscle tissue matrices have been decellularized, characterized in vitro, and validated in animal models [62–64]. We believe that the intact ECM proteins play an important role in tissue maintenance and reconstruction of skeletal muscle tissue; furthermore, they modulate cell adhesion and migration, growth factor storage and release, and satellite cell activation and differentiation [51, 52]. Moreover, immune responses to allogeneic or xenogeneic cell sources and tissues involved the recognition of major histocompatibility complex (MHC) molecules expressed on the surface of the cells by the immune cells of the host. The decellularization process removes all cellular components, including DNA, and leaves a structure comprised of ECM materials. Thus, minimal immune rejection response may occur with the transplantation of these allogeneic or xenogeneic nature scaffolds .
The use of native heart ECM is of special interest for cardiac-related cells due to the presence of intrinsic regulatory factors for cardiac function [65–67]. To engineer a tissue construct, cells are dependent on the ECM providing tissue-specific molecular, structural, and mechanical signals that regulate cell behaviors. Thus, scaffold materials for engineered cardiac tissue should mimic the native heart ECM in order to provide a natural microenvironment for the cardiac cells. Recently, incorporation of heart ECM into a bioink formulation for bioprinting cardiac tissue has been suggested . In an attempt to produce such a unique ECM component, decellularization technique of native tissues has been a very attractive method, while natural polymers and synthetic polymers are unable to replicate the precise spatial organization of complex structures like cardiac tissue . Generally, decellularization strategy is to remove cellular components from tissues via mechanical or chemical manipulation to produce ECM-rich matrices.
The high degree of evolutionary preservation of ECM components allows the use of acellular matrices derived from xenogeneic sources. Various decellularized tissue matrices have been utilized successfully for tissue engineering in animal models, and several xenogeneic products have received regulatory approval for clinical use . These decellularized xenogeneic medical products are being introduced into the market. Despite many advantages, there are concerns about the use of decellularized xenogeneic tissue matrices. These include the potential for immunogenicity, the possible presence of infectious agents, variability among preparations, and the inability to completely specify and characterize the bioactive molecules of the material. Many of the soluble bioactive factors in the tissue matrices can be lost during the decellularization process.
2.2.3 Synthetic Biodegradable Polymers
Synthetic polymers are high molecular weight macromolecules consisted of covalently bound repeating units (monomers). These repeating units can react with each other to form high molecular weight homopolymers by mainly addition polymerization, step-growth polymerization, and ring-opening polymerization. In tissue engineering, biodegradable synthetic polymers offer a number of benefits as scaffolds or drug delivery vehicles [71, 72]. These polymers can be synthesized with reproducible quality controls and high purity and fabricated into various shapes with anticipated bulk and surface features. Specific advantages include the ability to tailor the mechanical properties and degradation kinetics of these polymeric materials to suit various tissue engineering applications. Linear aliphatic poly(α-hydroxy acids), such as PGA, PLA, and their copolymer [poly(lactide-co-glycolide) (PLGA)], are the most commonly used biodegradable polymers in tissue engineering (Fig. 2.1) . These biodegradable polymers have gained popularity due to their processing reliability, tunable mechanical properties, and controlled biodegradability, and they are Food and Drug Administration (FDA)-approved for human use in a variety of medical applications, including suture material, implant, and drug delivery systems.
Chemical structure of the commonly used biodegradable synthetic polymers in tissue engineering
The ester bonds in these polymers degrade by nonenzymatic hydrolysis, and their nontoxic degradation products are eliminated from the body in the form of carbon dioxide and water. The degradation rate of these polymers can be controlled by alteration of their crystallinity, molecular weight, and the copolymer ratio of monomers. Furthermore, the degradation rates that can be achieved range from several weeks to several months. Because these polymers are thermoplastics, they can be configured into a three-dimensional (3D) structure with a desired microarchitecture, shape, and dimension. However, synthetic polymers generally lack intrinsic biological properties, and their degradation products may cause adverse effects or alter local microenvironment in vivo. In addition, the surface hydrophobicity of synthetic polymers may mediate protein denaturation in the vicinity of the implant and induce fibrous encapsulation [73, 74].
Several research teams have explored the synthesis of novel polymeric biomaterials that unite the advantages of “smart” polymeric systems with the biological activities of proteins at the chemical level. The concept of these smart polymeric systems was initially derived from the development of polymeric materials that show large conformational changes in response to microenvironmental stimuli, such as temperature, ionic strength, pH, or light . The responses of the polymers may include precipitation or gelation, reversible adsorption on a surface, collapse of a hydrogel or surface grafting, and alternation between hydrophilic and hydrophobic states . In most of cases, the change in the state of the polymer is reversible. Beyond the physical properties of these polymers, imparting smart biomaterials with the specific properties of signaling molecules, such as ECM components and growth factors, can be achieved. Hence, smart polymeric biomaterials aim to actively participate in the biological system in the body for accelerating tissue maturation and, eventually, regeneration.
To maximize the role of biomaterials in tissue regeneration, a combination of two or more materials with unique properties is often selected for use as a scaffold. These hybrid biomaterial systems that possess biological and physiochemical functionality have been used. The main strategy of this hybrid biomaterial system is to achieve synergy between the desirable biomechanical properties of one component with the biological compatibility and physiological relevance of the other component . For example, natural polymers possess microstructures that contain chemically and structurally defined components. Combination of these materials with synthetic polymers would offer a convenient bridge between chemical and biosynthetic approaches, often within the context of thermodynamically driven assembly processes .
2.3 Basic Requirements of Biomaterials in Tissue Engineering
Biomaterials replicate the biologic and biomechanical features of the native ECMs found in tissues in the body by serving as a synthetic microenvironment. Biomaterials provide a 3D space for the cells to form new tissues with tissue-specific structure and function; moreover, they could allow for the delivery of cells and bioactive molecules (such as cell adhesive proteins and growth factors), to the targeted sites in the body . As the majority of mammalian cell types are anchorage-dependent, biomaterials provide a cell adhesion substrate that can deliver cells to specific sites with high loading efficiency and homogeneity. Biomaterials can also provide biomechanical support against in vivo forces to maintain the predefined 3D tissue structure during the neo-tissue formation. Also, bioactive molecules, including adhesive proteins and growth factors, can be loaded along with cells to regulate the cellular behaviors.
In order to minimize the host responses, including inflammation, biomaterials should be biodegradable and bioresorbable. Incompatible materials produce an inflammatory or foreign-body reaction that eventually leads to rejection and/or tissue necrosis. Degradation products, if produced, should be removed from the body through the metabolic pathway at a designed rate that keeps the concentration of these degradation products in the body at a tolerable level . Biomaterials should also provide a tissue-specific microenvironment in which appropriate regulation of cell behaviors (adhesion, proliferation, migration, and differentiation) can occur so that functional neo-tissue can be formed. Cell behaviors in the newly formed tissue have been shown to be regulated by multiple interactions of the cells with their microenvironment, including interactions with cell adhesion ligands  and with soluble growth factors . Since biomaterials provide temporary mechanical support while the cells undergo spatial tissue reorganization, the properly chosen biomaterials should allow the engineered tissue to maintain sufficient biomechanical integrity to support itself in early developmental stage. In late development, biomaterials begin degradation; hence tissue growth is unhindered .
2.3.1 Scaffolding System: Template
Biomaterials for cell-based tissue engineering approaches should be biodegradable and absorbable without eliciting inflammatory responses that interfere with cellular function and neo-tissue formation; thus it has been considered to their biological properties, biodegradability, and structural and mechanical properties. In tissue engineering, biomaterials should  facilitate the localization and delivery of somatic cells to specific sites in the body,  maintain a 3D architecture that permits the formation of neo-tissues, and  guide the development of neo-tissues with tissue-specific functions . In details, biomaterials can provide an adhesion substrate for transplanted cells and serve as a delivery vehicle into the sites of interest. A large surface area-to-volume ratio is desirable in order to allow for accommodation of a high number of cells, which dictates the use of porous scaffolds. It is often desirable for neo-tissues to have a predefined structure, which can be achieved by an appropriate selection of biomaterials with controlled degradable rates. In addition, biomaterials can provide a temporary mechanical support that is sufficient to withstand in vivo forces and maintain a spatial structure until neo-tissue is able to support itself. This is essential for hard, weight-bearing tissue regeneration like the bone and cartilage. The stability of biomaterials depends primarily on physical properties and chemical degradation . The mechanical properties of biomaterials can be adjusted to suit the needs of the target tissue or organ along with the fabrication technologies.
A porous architecture with interconnectivity that allows the cellular ingrowth and maximizes the oxygen and nutrient diffusion is considered as an essential scaffold fabrication requirement. Configuration of such a structure may require tissue-specific considerations, which include variations in pore morphology and surface characteristics. In addition, biomaterial should be biodegradable or amenable to long-term integration with the host tissue. Biodegradability allows for the gradual and orderly replacement of the scaffold with functional tissue and prevents the development of adverse chronic responses to the artificial tissue structure . Therefore, biomaterials should be selected based on an understanding of target tissues or organs biologically, anatomically, pathologically, and biomechanically.
The mechanism of biodegradation of synthetic polymeric biomaterials is mainly processed through passive hydrolysis, with the monomeric acids as the final metabolized product . The degradation rates are closely related to molecular weight, degree of crystallization, hydrophilicity/hydrophobicity, and surface-to-volume ratio. For instance, PGA is generally available in a highly crystalline form, and it possesses a relatively high melting point (~230 °C) and low solubility in organic solvents. However, its hydrophilic nature allows for rapid absorption of water leading to a fast degradation within a few weeks. In contrast, PLA is available in semicrystalline or amorphous forms, depending on the types of the d and l optical isomers of lactic acid present in the molecule. D-PLA and L-PLA are both highly crystalline, whereas the racemic copolymer D,L-PLA is completely amorphous. The presence of a methyl group makes PLA more hydrophobic than PGA, resulting in lower water absorption rates and correspondingly slower degradation, in the order of months . Moreover, the factors affecting biodegradation include chemical and structural configuration, fabrication parameters, sterilization methods, and implantation sites. For example, sterilization using γ-irradiation can decrease the polymer’s Mw by 30–40%. The irradiated polymers continue to decrease in Mw during storage at room temperature. The decline in Mw affects the mechanical properties and the biodegradation rate of the polymers [8, 11].
The degradation of polymeric biomaterials results in erosion (loss of mass), when the degradation products diffuse or dissolve into the body fluid. Polymer erosion is generally classified by bulk erosion and surface erosion . The polymer chain scission by the bulk erosion is wholly occurred, while the surface erosion is only occurred in the surface of the polymeric scaffolds. In the bulk erosion, the molecular weight and mechanical properties of the polymeric scaffolds decrease in time; however, the external dimension of the scaffolds remains essentially unchanged until the scaffold is fully degraded. In the surface erosion, the external dimension and mass of the scaffolds decrease in time.
Unlike the synthetic polymers, the biodegradation of the natural polymers is mainly dependent on the enzymatic reaction. This is the microenvironmental responsiveness of the natural polymers through degradation and remodeling by cell-secreted enzymes . The biodegradation rate of these natural polymers can be controlled by various cross-linking processes. For instance, a collagen scaffold that is designed for a long-term in vivo use would require a cross-linking process to improve its resistance to the proteolytic degradation and hydrolysis and to enhance its mechanical strength . Glutaraldehyde is a commonly used cross-linker for tissue matrices because it has amine-reactive homobifunctional groups . Carbodiimide chemistry has been used extensively to cross-link collagen scaffolds and to immobilize other proteins to the scaffolds. Particularly, 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC) has been used as a zero-length cross-linker that directly binds between collagen molecules .
Biomaterials used for cell-based tissue engineering serve as an artificial ECM and bring forth biological, biochemical, and biomechanical functionalities of native ECM found in tissues and organs. Therefore, cell-biomaterial interactions, such as cellular adhesion on the matrix surface, cell maturation, ECM production, and proliferation, are critical for the success of this strategy. ECM is known to affect cellular differentiation and function; thus, biomaterials can be designed to deliver cells and bioactive molecules (e.g., cytokines and growth factors) to a designated region for neo-tissue formation with enhanced functionality [4–6].
2.4 Advanced Biomaterial Systems in Tissue Engineering
Future advances in biomaterials for tissue engineering rely on the development of the biomimetic materials that can utilize in vivo microenvironment for efficient tissue regeneration and better integration with host tissue (Fig. 2.2). Desirable biomimetic materials could be designed as a smart scaffolding system that mimics the natural biological system and integrates the necessary structural and biological properties. Thorough understanding of biomaterials science combined with extensive knowledge of the clinical challenges and cell biology is vital for development of novel advanced biomaterials to be used for tissue engineering applications.
Schematic illustration of interactions between endogenous stem cells and biomimetic materials. Stem cell fate in a particular microenvironment is regulated by intricate reciprocal molecular interactions with its surroundings
2.4.1 Surface Modification
One commonly used technique is the surface modification of polymeric biomaterials. Chemical and physical modification of the biomaterials surface can increase their biological properties (Fig. 2.3). Chemical procedures, such as oxidation, hydrolysis, and quaternization, change the surface chemistry and functionality of biomaterials. Polymerization or grafting of water-soluble polymers to a biomaterial surface is usually performed to improve the protein adsorption and further cell adhesion [86, 87]. In addition, bioactive molecules, such as enzymes, drugs, proteins, peptide sequences, and antibodies, have been immobilized onto biomaterial surfaces to achieve added functionality . These approaches provide an opportunity to tailor a biomaterial surface so that it has specific affinities, site recognition properties, and controlled mobility, while the bulk properties of biomaterials are still maintained.
Schematic diagram of surface modification methods of Polymeric biomaterials
2.4.2 Drug Delivery System
Biomaterial functionalization has also allowed enhanced cellular interactions via delivery of bioactive molecules from an implanted scaffold . Bioactive molecules, such as cytokines and growth factors, are powerful regulators of biological functions, which include the cell migration, proliferation, and differentiation. Incorporation of bioactive molecules into biomaterial scaffolds is a technique to improve the outcome of cell-based tissue engineering. To employ this technique, an understanding of the biological activities of molecules is necessary. For example, the biological activity of growth factors is dependent not only on their presence in solution but also on their interactions with the surrounding ECM. Some growth factors are most effective when released over a prolonged period of time, whereas others are more effective when delivered in a bolus. Some factors are active while tethered to a material, whereas others are only active when they have been released from the biomaterial and are internalized into a cell. These considerations must be taken into account when designing a delivery system .
2.4.3 ECM-Mimetic Biomaterials
Development of advanced biomaterials has been extended to the design, discovery, and evaluation of biologically active materials. This required the relatively straightforward chemical synthesis of new polymeric materials, coupled with a search for novel activities and examination of their performance in biological systems. By adapting the combinatorial library approach that has been well established for synthetic peptides and small molecule drugs with novel high-throughput assays, thousands of candidate chemicals can be synthesized and tested. As one example, screening process 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 (ESCs) in unexpected ways . More revolutionary developments in biomaterials science will continue to arise as tissue engineering is combined with technology-based approaches. Basic understanding of the 3D structure of existing biological molecules is being applied to a “bottom-up” approach to generate new, self-assembling supramolecular architectures .
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 microenvironmental cues are inherent in some of these biomaterials . Recent advances include the design of short peptides (e.g., heptamers) based on coiled-coil motifs that reversibly assemble into nanofilaments and nanoropes, without excessive aggregation . These novel peptide amphiphiles can be induced to self-assemble by changes in concentration, pH, or the level of divalent cations . Branched structures can be designed to present bioactive sequences such as arginine-glycine-aspartic acid (RGD) to cells by nanofibrous hydrogels or as coatings on typical tissue engineering scaffolds . In addition, assembly can occur under conditions that permit the entrapment of cells in the resulting nanofibrous matrix . The entrapped cells retain motility and the ability to proliferate. Peptide-based nanofibers can be designed to present bioactive sequences to cells at very high density, substantially exceeding that of corresponding peptide epitopes in biological ECM.
A pentapeptide epitope of laminin, isoleucine-lysine-valine-alanine-valine (IKVAV), known to promote neurite extension, was incorporated into peptide amphiphiles capable of self-assembly into nanofibers that form a hydrogel . When neural progenitor cells (NPCs) capable of differentiating into neurons or glial cells were encapsulated during assembly of the nanofibers, they survived over a few weeks in culture. Moreover, even without the addition of neurotrophic factors, they displayed neuronal differentiation as exemplified by the extension of large neurites that were obvious after 1 day in culture and by their expression of β-tubulin III. The production of neuron-like cells from neural progenitors, whether dissociated or grown as clustered neurospheres, was more rapid and robust in the IKVAV-containing peptide amphiphile (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 hydrogels 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 tissue engineering and regenerative medicine.
Proteins have been conjugated either randomly or in a site-specific manner, through engineering of protein to introduce a reactive amino acid at a particular site . If a conjugation site is introduced near the ligand-binding domain of protein, it has been shown that induction of a change in conformational state of the novel polymers can serve to regulate the protein’s activity [100–103]. This may allow selective capture and recovery of specific cell types, delivery of cells to a desired location, and modulation of enzymes that influence the tissue remodeling process. The discovery of new classes of biomaterials may provide yet another opportunity to overcome the current limitations in tissue engineering.
2.4.4 Biophysical Signals
Cells are exposed to a tissue-specific microenvironment in vivo, and they respond to numerous chemical and physical stimuli. As a result, cellular functions like cell adhesion and proliferation, protein synthesis, and cytoskeletal architecture are critically influenced by the microenvironment of the cells. To maintain the phenotype expression and differentiation of tissue-specific cells, numerous cues have been applied to biomaterial scaffolds to mimic natural ECM microenvironments. Cells can be directly influenced by topographical cues of their surrounding microenvironment. The topography of the biomaterial surface can influence the cell adhesion and proliferation, which further affects cellular functions. It has been determined that natural ECM has a nanoscaled fibrous structure that supports the cellular interactions and stimulates the cellular functions. This substrate nanoscale topography provides a tool for investigating complex cell functions, such as adhesion, migration, cytoskeleton reorganization, and cell polarization [104, 105].
Physical contact of cells on the biomaterial surface is essential to anchorage dependence and function. Current understanding reveals that stem cell fates can be significantly affected by the substrate elasticity from the surrounding microenvironment, while stem cell differentiation induced by soluble biological signals has been widely developed. For example, neural cell differentiation can be induced on soft matrices (high elasticity), which mimic the mechanical properties of the brain tissue, while myogenic and osteogenic differentiation require stiffer matrices (low elasticity) that mimic muscle and collagenous bone tissues. Soft matrices that mimic brain are neurogenic, stiffer matrices that mimic muscle are myogenic, and comparatively rigid matrices that mimic collagenous bone prove osteogenic . These studies have demonstrated that specific stem cell differentiation can be achieved using substrates or matrices with particular surface elasticity/stiffness. Modification of the cross-linking degree of polymeric hydrogels is the basis of substrate-controlled elasticity.
2.5 Scaffold Fabrication Technologies
Biomaterials used for tissue engineering applications require different spatial configurations, depending on the desired outcome. In tissue engineering, biomaterials need to be fabricated to facilitate regeneration of the target tissues or organs. Thus, starting with simple foams and fibers, the porosity and pore interconnectivity of biomaterial scaffolds has quickly become a central theme, with pore dimensions and physical properties being vital in promoting cell seeding, migration, proliferation, and de novo production of ECM molecules. Many fabrication technologies have been employed to configure biomaterials in this way for tissue-engineered scaffolds. In such instances, biomaterials provide the structural support for cells to proliferate and maintain their function. To permit cell accommodation, biomaterials must possess adequate surface area for cell attachment and distribution. This is usually achieved by creating interconnecting pores that allow for maximum diffusion of oxygen and nutrients. Thus, numerous fabrication technologies have been developed to create porous scaffolds, which include solvent casting/particulate leaching, freeze-drying, fiber bonding, electrospinning, phase separation, gas foaming, high pressure processing, melt molding, membrane lamination, hydrocarbon templating, cell printing, and solid free-form fabrication (SFF) methods.
Hydrogels have been used in injectable approaches for cell therapy and tissue engineering, which offer many advantages. These include the repair of tissue defects or injection of cells for achieving cellular function . However, while injection therapy is simple and does not require invasive surgical procedures, its application is limited to areas of the body where injection is feasible. Injectable biomaterials usually take the form of liquids or gels, which can be configured from synthetic and natural polymers. A “hydrogel” is defined as a 3D network swollen by water which is usually the major component of the hydrogel system. Hydrogels are formed by cross-linking water-soluble polymers to form an insoluble, hydrophilic polymer network. These polymer networks are capable of absorbing water from a small fraction of their dry weight up to thousands of times their dry weight . Representative hydrogels include poly(ethylene glycol) (PEG), poly(ethylene oxide) (PEO), poly(vinyl alcohol) (PVA), poly(acrylic acid) (PAA), Pluronics (PEO-PPO-PEO copolymers), agarose, alginate, chitosan, collagen, fibrin, gelatin, and hyaluronic acid (HA).
In the past decade, progress in creating 3D cellular microenvironments using hydrogels has appeared to have similar elasticity to that of native tissues. The structure and composition of hydrogels can be tailored to bear the suitable chemical, biological, and physical cues that encourage the development of tissue constructs . For example, matrix metalloproteinase (MMP)-sensitive PEG-based hydrogels have been developed to direct differentiation of pluripotent cardioprogenitor cells using P19 embryonal carcinoma cells . In order to systematically probe 3D microenvironmental effects on P19 embryonal carcinoma cell differentiation, matrix elasticity, MMP-sensitivity, and the concentration of a matrix-bound RGDSP peptide were modulated. Results indicated that soft PEG-based hydrogels mimicking the elasticity of embryonic cardiac tissue could direct cardiac commitment of pluripotent P19 cells.
The most common cell encapsulation system involves alginate microcapsules, which can enclose living cells or other molecules. Cell encapsulation techniques such as this allow allogeneic and xenogeneic cell sources to be considered for clinical translation, and this is important as autologous cells can often be limited in supply. In order to use allogeneic or xenogeneic cell sources, the immune responses of the host need to be addressed. An immunoisolation technique using cell encapsulation permits the use of these cell sources without the use of immunosuppressants. Cell microencapsulation can be used for cell therapy as well as for delivering biological products. This technique consists of enclosing biologically active substances within a polymeric hydrogel surrounded by a semipermeable membrane that is designed to prevent immune recognition and subsequent rejection. Ideally, the encapsulating membrane allows the bidirectional diffusion of nutrients, oxygen, and waste as well as the secretion of therapeutic products while preventing immune cells and antibodies, which would normally destroy the enclosed cells, from entering the capsule . Pancreatic islets, hepatocytes, renal cells, parathyroid cells, and adrenal chromaffin cells have been immunoisolated for cell-based therapeutic delivery systems in this manner. Several polymeric encapsulation systems have been developed and are currently being tested in clinical trials [111, 112].
While many other scaffold fabrication methods are used in tissue engineering applications, few can provide scaffolds with critical similarities to natural ECM that electrospinning of nanofibers can provide. Other fabrication processes allow for the inclusion of natural polymers such as collagen into the scaffold, but some, such as melt molding, require high temperatures which may denature proteins. Others, such as solvent casting and particulate leaching, provide scaffolds with very high porosity but potentially low interconnectivity and may not allow for easy shaping . Electrospinning has become a popular alternative fabrication method, as it can be applied to many disciplines and is a relatively simple and inexpensive scaffold fabrication process. It relies on a high-voltage direct or alternating current source to charge a polymer solution or melt contained in a syringe . Production of nanofibers by electrospinning provides excellent pore interconnectivity and precise control of porosity through adjustment of fiber size . This interconnectivity can allow for integration of cells into the scaffold if the pore size is large enough  and also allows for dissolution of soluble factors and nutrients throughout the scaffold. These properties closely mimic the natural properties of ECM, and the parameters of the electrospinning process can be controlled so as to adjust to features of targeted tissue types. For instance, arteries are composed of a layer called the intima, which consists of an endothelial cell (EC) layer lining the lumen of the artery, which sits on a basal lamina. Surrounding this layer is the tunica media, which is a layer of connective tissue and smooth muscle cells (SMCs). To mimic this vessel structure, our lab showed that it is possible to create a bilayered vascular scaffold which supports EC growth on an electrospun intimal layer and SMC growth on an electrospun media layer (Fig. 2.4) .
Bilayered electrospun PCL/collagen vascular scaffolds; (a) gross appearance, (b) fluorescent images of bilayered scaffolds incorporated fluorescent dyes. SEM images of cross-sectional (c) outer layer, (d) entire, and (e) interface between outer and inner layers of bilayered scaffolds. Reproduced with permission from Ju et al. 
Electrospun fibrous scaffolds can also be functionalized by adding biochemical and biomechanical cues to enhance cellular interactions; however, specific interactions between cells and electrospun scaffolds functionalized through surface modifications and bioactive factor incorporation are still poorly understood. A better understanding of the specific cues that enhance cell adhesion, proliferation, and guidance of cells seeded on a scaffold, as well as the cues that could affect host cell infiltration, differentiation, and vascularization both in vitro and in vivo, is crucial for the advancement of tissue engineering applications of electrospinning.
2.5.3 Computer-Aided Scaffold Fabrication
For clinical translation, a fabrication method should provide a scaffold with appropriate shape and size to reconstruct damaged tissues in patients. Computer-aided fabrication or additive manufacturing of tissue-engineered constructs is a rapidly emerging technology that promises to accelerate the clinical translation of many tissue engineering applications. This technology is based on the ability to build tissue-engineered constructs with highly reproducible architecture and to provide compositional variation across the entire scaffold through tightly controlled, computer-driven fabrication. To accomplish this, a conventional inkjet printer can be used [117–120]. This thermal inkjet printer produces small air bubbles consisting of cells and a hydrogel by heating, and then bubbles are collapsed to provide the pressure pulse to eject a very tiny drop of “bioink” out of the nozzle. This printing technique uses aqueous processes that have been shown to induce little damage to cells. In the extrusion printing method, a pneumatic dispensing system has been utilized to build various 3D architectures by depositing a continuous stream of cell-laden hydrogel . These approaches take advantage of rapid prototyping assisted by computer-assisted design (CAD) procedures to build a tissue construct with geometric control of its internal structure and external shape. We have recently showed that viable 3D heterogeneous construct with multiple cell types could be created by printing the cells with hydrogel layer-by-layer . This cell-printed constructs were able to survive and mature into tissues with adequate vascularization when implanted in vivo (Fig. 2.5) .
3D multiple cell positioning: (a–c) type I pattern and (d–f) type II pattern. (a, d) Two types of 3D patterning, including cell-A (red), cell-B (blue), and PCL (green), were fabricated. (b, e) Photographs and (c, f) fluorescent image of the 3D printed patterns. (g) Cell viability (%) was over 95% on day 0 and then maintained on days 3 and 6 (n = 3). (h) Cell proliferation results showed that the number of cells continuously increased over a 15-day period, and no significant differences between the control and the printed constructs were observed (n = 5). Reproduced with permission from Kang et al. 
In 2016, our laboratory has reported a novel integrated tissue and organ printing (ITOP) system that makes it possible to fabricate a high strength tissue construct and precisely place various cell types within a single architecture [122, 123]. This system consists of multiple nozzles that allow delivery of multiple types of biomaterials and cells. The goal of ITOP is to concurrently deliver a supportive 3D template (mostly biodegradable thermoplastic polymers) and 3D–patterned deposition of cell-laden hydrogel biomaterials in a precise manner. This approach is borne out of the challenges of tissue engineering 3D tissue constructs.
2.6 Tissue Engineering Applications
Several therapeutic approaches on the cellular level have been investigated for direct treatment of medical conditions or for use of implanted engineered tissue constructs. A variety of tissue engineering products have been tested by injection, gene delivery, implantation, encapsulation, or combined systems. Applications in tissue engineering of the several therapies are described.
2.6.1 Bulking Agents for Urinary Incontinence and Vesicoureteral Reflux
Injectable bulking agents have been used endoscopically in the treatment of both urinary incontinence and vesicoureteral reflux. The advantages in treating urinary incontinence and vesicoureteral reflux with this minimally invasive approach include the simplicity of a quick outpatient procedure and the low morbidity associated with it. Investigators are seeking alternative biomaterials that would be safe for human use. The ideal substance for the endoscopic treatment of reflux and incontinence should be injectable, nonantigenic, nonmigratory, volume stable, and safe for human use. Toward this goal, long-term studies were conducted to determine the effect of injectable chondrocytes in vivo. It was initially determined that alginate, a liquid solution suspended with chondrocytes, could serve as a synthetic substrate for the injectable delivery and maintenance of cartilage architecture in vivo. This system was adapted for the treatment of vesicoureteral reflux in a mouse model  and a porcine model . The use of autologous cartilage for the treatment of vesicoureteral reflux in humans would satisfy all the requirements for an injectable substance.
Two multicenter clinical trials were conducted using this technology. Patients with vesicoureteral reflux were treated at ten centers throughout the USA. The patients had a similar success rate as with other injectable substances, in terms of cure. Chondrocyte formation was not noted in patients who had treatment failure. It is supposed that the patients who were cured have a biocompatible region of engineered autologous tissue present, rather than a foreign material . Patients with urinary incontinence were also treated endoscopically with injected chondrocytes at three different medical centers. Phase 1 trials showed an approximate success rate of 80% at follow-up 3 and 12 months postoperatively . Therefore, human application of cell-based tissue engineering for urologic applications occurred successfully with the injection of chondrocytes for the correction of vesicoureteral reflux in children and for urinary incontinence in adults [126, 127].
2.6.2 Injection Therapy Using Muscle Progenitor Cells
The potential use of injectable cultured primary muscle progenitor cells (MPCs) for the treatment of stress urinary incontinence has been investigated . Fluorescence-labeled MPCs were injected directly into the proximal urethra and lateral bladder walls of nude mice with a microsyringe in an open surgical procedure. Tissue harvested up to 35 days after injection contained the labeled MPCs, as well as evidence of differentiation of the labeled MPCs into newly formed myofibers. It was reported that a significant portion of the injected MPC population persisted in vivo. Similar techniques of sphincteric-derived MPCs have been used for the treatment of urinary incontinence in a clinical trial . The fact that MPCs can be labeled, survive after injection, and begin the process of myogenic differentiation further supports the feasibility of using cultured cells of muscular origin as an injectable bioimplant.
The use of injectable MPCs has also been investigated for use in the treatment of urinary incontinence due to irreversible urethral sphincter injury or maldevelopment. MPCs are the quiescent satellite cells found in each myofiber that proliferate to form myoblasts and, eventually, myotubes and new muscle tissue. Previously, intrinsic MPCs have been shown to play an active role in the regeneration of injured striated urethral sphincter. In a subsequent study, autologous MPCs were injected into a rat model  and a canine model [10, 131] of urethral sphincter injury, and both replacement of mature myotubes, as well as restoration of functional motor units, were noted in the regenerating sphincteric muscle tissue. This was the first demonstration of the replacement of both sphincter muscle tissue and its innervation by the injection of MPCs. As a result, MPCs may be a minimally invasive solution for urinary incontinence in patients with irreversible urinary sphincter muscle insufficiency.
2.6.3 Endocrine Replacement
Patients, who have a testicular dysfunction and hypogonadal disorders, are dependent on androgen replacement therapy to restore and maintain physiological levels of serum testosterone and its metabolites, dihydrotestosterone, and estradiol . Currently available androgen replacement modalities, such as testosterone tablets and capsules, depot injections, and skin patches, may be associated with fluctuating serum levels and complications such as fluid and nitrogen retention, erythropoiesis, hypertension, and bone density changes . Since Leydig cells of the testes are the major source of testosterone in men, implantation of heterologous Leydig cells or gonadal tissue fragments has previously been proposed as a method for chronic testosterone replacement . However, these approaches were limited by the failure of the tissues and cells to produce testosterone.
Encapsulation of cells in biocompatible and semipermeable polymeric membranes has been an effective method to protect against host immune responses while maintaining viability of the cells and allowing the secretion of desired therapeutic agents [135, 136]. Encapsulation of Leydig cells in alginate-poly(l-lysine) was used as a novel method for testosterone delivery in vivo . Elevated stable serum testosterone levels were noted in castrated adult rats over the course of the study, which suggests microencapsulated Leydig cells may be a potential therapeutic modality for testosterone supplementation.
2.6.4 Urethral Tissue
Various approaches have been proposed for the regeneration of urethral tissue. PGA nonwoven meshes without cells [137, 138] or with cells  were used to regenerate urethra tissues in various animal models. Acellular tissue matrices such as decellularized bladder submucosa  and urethral submucosa  have also been tested in various animal models for urethral reconstruction. Bladder submucosa matrix (BSM) proved to be a suitable graft for repair of urethral defects in rabbits . The grafts demonstrated a normal urothelial luminal lining and organized muscle bundles on the outer portion of the graft. These results were confirmed clinically in a series of patients with a history of failed hypospadias reconstruction. The urethral defects in these patients were repaired with human BSMs . An engineered urethral tissue was created by anastomosing the matrix to the urethral plate in an onlay fashion. The size of the engineered urethra tissue ranged from 5 to 15 cm. At the 3 years follow-up, three of the four patients had a successful outcome in regard to cosmetic appearance and function. One patient, who had a 15 cm neo-urethra tissue formed, developed a subglanular fistula. Similar results were obtained in pediatric and adult patients with primary urethral stricture disease using the same BSMs .
Another study in 30 patients with recurrent stricture disease showed that a healthy urethral bed (two or fewer prior urethral surgeries) was needed for successful urethral reconstruction using the BSM grafts . More than 200 pediatric and adult patients with urethral disease have been successfully treated in an onlay manner with BSMs. One advantage of this acellular scaffold over the traditional nongenital tissue grafts that have been used for urethroplasty is that the acellular scaffold is “off the shelf.” Eliminating the necessity of additional surgical procedures for graft harvesting may decrease operative time as well as the potential morbidity due to the harvest procedure.
These techniques, which employed acellular tissue scaffolds that had not been reseeded with cells, were applied experimentally and clinically in a successful manner for onlay urethral repairs. Further study, however, indicated that when tubularized urethral repairs with unseeded scaffolds were attempted experimentally, adequate urethral tissue regeneration was not achieved, and complications, such as graft contracture and stricture formation, occurred . To determine if seeding the scaffold with cells from the urinary tract could improve the results of tubularized urethral repairs, autologous rabbit bladder epithelial cells and SMCs were grown and seeded onto tubular scaffolds. Entire urethra segments were then resected, and urethroplasties were performed with tubularized scaffolds either seeded with cells or without cells. The tubularized scaffolds seeded with autologous cells formed neo-tissue which was histologically similar to native urethra tissue. The tubularized scaffolds without cells lead to poor tissue development, fibrosis, and stricture formation. These findings were confirmed clinically when a trial using tubularized non-seeded scaffolds for urethral stricture repair was performed in eight evaluable patients. Two patients with short inflammatory strictures maintained urethral patency. Stricture recurrence developed in the other six patients within 3 months of surgery .
In 2011, Raya-Rivera et al. were able to show that synthetic polymers can also be used in urethral reconstruction when they are tubularized and seeded with autologous cells . This group used PGA-PLGA scaffolds seeded with autologous cells derived from bladder biopsies taken from each patient. These cell-seeded scaffolds were then used to repair urethral defects in five boys. Upon follow-up evaluation, it was found that most of the boys had excellent urinary flow rates postoperatively, and voiding cystourethrograms indicated that these patients maintained wide urethral calibers. Urethral biopsies revealed that the grafts had developed a normal appearing architecture consisting of urothelial and muscular tissues (Fig. 2.6).
Neo-urethra implantation and clinical outcomes; (a) a cell-seeded graft sutured to the normal urethral margins. Scanning electron microscopy of seeded scaffolds 6 days after seeding; (b) urothelial cells, (c) smooth muscle cells, and (d) scaffold without cells. Voiding cystourethrograms of a patient (e) before surgery (arrows show the abnormal margins), (f) 12 months after surgery (arrows show margins of tissue-engineered urethras), and (g) at last follow-up. Histological and immunohistochemical analyses; (h) H&E, (i) α-actin, (j) desmin, (k) mvosln, and (l) AE1/AE3. Reproduced with permission from Raya-Rivera et al. 
The challenges in the bladder augmentation are often associated with renal function and quality of life . Gastrointestinal segments are most commonly used for this purpose in the clinic, but they are designed to absorb specific solutes, whereas bladder tissue is designed for the excretion of solutes. Numerous studies have attempted alternative biomaterials and tissue-engineered constructs for bladder replacement or repair due to the problems encountered with the use of gastrointestinal segments. The ability to use donor tissue efficiently and to provide the right conditions for long-term survival, differentiation, and growth determines the success of cell transplantation strategies for bladder reconstruction. These principles were applied in the creation of tissue-engineered bladders in an animal model that required a subtotal cystectomy with subsequent replacement with a tissue-engineered organ in beagle dogs . Bladder urothelial cells and SMCs were separately expanded from an autologous bladder biopsy and seeded onto a bladder-shaped biodegradable polymeric scaffold. The results from this study revealed the possibility of engineering bladder constructs that are anatomically and functionally normal.
Even though the clinical translation of tissue engineering and regenerative medicine is slowly progressing, the success of tissue engineering for some applications was already demonstrated clinically. A clinical experience involving engineered bladder tissue for cystoplasty reconstruction was conducted starting in 1999. A small pilot study of seven patients was reported, using a collagen scaffold seeded with cells either with or without omentum coverage, or a combined PGA/collagen scaffold seeded with cells and omental coverage. Patient’s own bladder cells were isolated and expanded in vitro prior to seeding on a scaffolding system of collagen and PGA. The scaffold was designed to replace the bladder size to improve its compliancy and patient continence (Fig. 2.7a–c). The tissue-engineered constructs were implanted for a period up to 61 months. Our outcomes show the definitive improvement of bladder compliance and capacity while restoring physiological function (Fig. 2.7d, e). In addition, the engineered bladder biopsies showed an adequate structural architecture and phenotype (Fig. 2.7f–k). Although the experience is promising in terms of showing that engineered tissues can be implanted safely, it is just a start in terms of accomplishing the goal of engineering fully functional bladders. Further experimental and clinical work is being conducted.
(a) Scaffold seeded with cells, (b) engineered bladder anastomosed to native bladder with running 4–0 PGA sutures, and (c) implant covered with fibrin glue and omentum. (d) Preoperative and (e) 10 months postoperative cystograms and urodynamic findings in patient with a collagen-PGA scaffold engineered bladder. (f–h) cystoscopic biopsies of implanted engineered bladders 31 months after augmentation shows extent of regeneration. Engineered bladder tissue showed tri-layered structure, consisting of lumen lined with urothelial cells, U, surrounded by submucosa, S, and muscle, M. (f) H&E, immunocytochemical analysis with (g) anti-pancytokeratin AE1/AE3 antibodies, and (h) anti-α smooth muscle actin antibodies showed presence of phenotypically normal urothelium and smooth muscle. (i–k) native bladder tissue. Original magnification: ×100. Reproduced with permission from Atala et al. 
2.6.6 Male and Female Reproductive Organs
Reconstructive surgery is required for a wide variety of pathologic penile conditions such as penile carcinoma, trauma, severe erectile dysfunction, and congenital conditions such as ambiguous genitalia, hypospadias, and epispadias. One of the major limitations of phallic reconstructive surgery is the scarcity of sufficient autologous tissue. The major components of the phallus are corporal SMCs and ECs. The creation of autologous functional and structural corporal tissue de novo would be beneficial. Autologous cavernosal SMCs and ECs were harvested, expanded, and seeded on decellularized tissue matrices and implanted in a rabbit model [149, 150]. The engineered corpora demonstrated structural and functional parameters similar to native tissue and male rabbits receiving the bilateral implants were able to successfully impregnate females. This study demonstrated that neo-corpora could be engineered for total pendular penile corporal body replacement . Such technology has considerable potential for patients requiring penile reconstruction.