Polysaccharide Based Scaffolds for Soft Tissue Engineering Applications

Soft tissue reconstructs require materials that form three-dimensional (3-D) structures supportive to cell proliferation and regenerative processes. Polysaccharides, due to their hydrophilicity, biocompatibility, biodegradability, abundance, and presence of derivatizable functional groups, are distinctive scaffold materials. Superior mechanical properties, physiological signaling, and tunable tissue response have been achieved through chemical modification of polysaccharides. Moreover, an appropriate formulation strategy enables spatial placement of the scaffold to a targeted site. With the advent of newer technologies, these preparations can be tailor-made for responding to alterations in temperature, pH, or other physiological stimuli. In this review, we discuss the developmental and biological aspects of scaffolds prepared from four polysaccharides, viz. alginic acid (ALG), chitosan (CHI), hyaluronic acid (HA), and dextran (DEX). Clinical studies on these scaffolds are also discussed.


Introduction
Soft tissues are complex fiber-reinforced structures, generally distinguishable from hard tissues by their high water content [1]. They are continuously invaded by trauma, invasive surgery, and aging. This often leads to impaired physiological functions, large scale tissue loss, and even organ failure [2]. The restorative approaches include direct administration of primary or genetically engineered cells of auto-, allo-, or xenogeneic origin [3,4], and transplantation of cells seeded into tissue-like three dimensional (3D) scaffolds. Devoid of a stiff matrix, the former approach is associated with serious obstacles, such as the rapid escape of cells, suboptimal dispersion, insufficient vascularisation, donor site morbidity, potent immunogenic response, and long-term administration of immunosuppressive agents [5][6][7][8][9][10]. The implantation of autologous cells is challenging due to difficulty in harvesting clinical-grade cells in sufficient number, especially in aged recipients or when the damage is high [11]. Moreover, cell harvesting requires a second surgical site. This two-stage procedure increases surgery time and patients may suffer nerve damage at the harvest site [12]. The instillation procedure via traditional hand held injections imposes a pronounced surgical stress on suspended cells [13]. Studies report that 80%-90% of transplanted cells die within the first 72 h of injection [14]. More importantly, cellular de-differentiation during in vitro propagation may alter the biosynthetic properties of autologous cells [15].
Seeding the lineage-and tissue-specific progenitors, derived from patient's normal tissue or donor, into scaffolds is a rapidly expanding tissue engineering (TE) alternative. In the 1980s, TE was understood as the application of prosthetic devices and surgical manipulation of tissues [16]. Despite these merits, the application of natural polysaccharides for scaffold preparation is associated with certain limitations. Their molecular weight distribution, branching, and sequence may not be consistent. In addition to influencing rheology, these variations may be detrimental to biorecognition events. Another noteworthy obstacle is the inferior mechanical strength of polysaccharide gels. It leads to quick hydrolysis and displacement of the formulation away from the injection site. For instance, deterioration in the viscoelasticity of hyaluronic acid occurs through the production of oligosaccharides and low molecular weight fragments [52]. Loss of viscoelasticity at physiological temperatures can be circumvented through age-old crosslinking methods [58,59]. Besides, gel strength has been improved through the incorporation of additives and/or polymers of a desirable molecular weight [60][61][62][63]. Gelatin is a good choice as a blend component (Table 2). With the presence of arginine, glycine, and aspartic acid (RGD) tripeptide in the backbone, gelatin acts as a fibroblast-attractant. Simultaneously, it promotes epithelialization and granulation tissue formation [64][65][66][67]. It has been shown to undergo proteolytic degradation without producing antigenic fragments [38].
In this review, we discuss the developmental and biological aspects of scaffolds prepared from four polysaccharides, viz. alginic acid (ALG), chitosan (CHI), hyaluronic acid (HA), and dextran (DEX). An emphasis is placed on scaffolds developed through physical/chemical modifications using crosslinking, grafting, polyion complexation, and blending ( Table 2). Clinical studies on these scaffolds are also covered (Table 3).  Figure 1. Schematic structure of polysaccharides. Active centre in the repeating unit of each polysaccharide is shown in red font.
Despite these merits, the application of natural polysaccharides for scaffold preparation is associated with certain limitations. Their molecular weight distribution, branching, and sequence may not be consistent. In addition to influencing rheology, these variations may be detrimental to biorecognition events. Another noteworthy obstacle is the inferior mechanical strength of polysaccharide gels. It leads to quick hydrolysis and displacement of the formulation away from the injection site. For instance, deterioration in the viscoelasticity of hyaluronic acid occurs through the production of oligosaccharides and low molecular weight fragments [52]. Loss of viscoelasticity at physiological temperatures can be circumvented through age-old crosslinking methods [58,59]. Besides, gel strength has been improved through the incorporation of additives and/or polymers of a desirable molecular weight [60][61][62][63]. Gelatin is a good choice as a blend component (Table 2). With the presence of arginine, glycine, and aspartic acid (RGD) tripeptide in the backbone, gelatin acts as a fibroblast-attractant. Simultaneously, it promotes epithelialization and granulation tissue formation [64][65][66][67]. It has been shown to undergo proteolytic degradation without producing antigenic fragments [38].

Chitosan
Chitosan (CHI) is obtained from the partial or full deacetylation of chitin (the second most abundant biopolymer after cellulose, found in the exoskeleton of crustaceans and endoskeleton of molluscs). The protonation of amine groups during dissolution imparts a positive charge, following which it quickly adheres to negatively charged substrate surfaces. Readers are referred to some earlier reviews on biochemistry and biomedical applications of CHI [78,79]. Depending on the source, chitin exists in αor βcrystallographic forms. As against to anti-parallel chain organization of α-form, β-chitin exhibits parallel organization. The latter configuration though allows limited probability of intermolecular hydrogen bonds, but it improves accessibility to chemical modification or deacetylation [80,81]. This is evident in the findings of Reys et al. [82], who investigated the influence of freezing temperatures (−80 and −196 • C) upon scaffold formation behavior of αand β-chitin, with a deacetylation degree (DD) of 76.6% and 91.2%, respectively. Although both the scaffolds exhibited stability against lysozyme up to 4 weeks, those prepared at −196 • C displayed a compact structure and smaller pores. The β-chitin scaffold presented similar morphological features and swelling profile, but superior mechanical properties attributable to its higher DD [82].
Tissue engineering applications of CHI emerge from the properties, such as hydrophilicity, polyelectrolyte behavior, mucoadhesion, hemostatic action, and structural similarity to native extracellular proteoglycans. Its polar groups and physicochemical properties provide a favorable non-protein environment for cell adhesion and proliferation [83]. It easily forms a blend with other polymers through electrostatic interactions and confers antimicrobial properties to the final composition [1,21,84].

Alginic Acid
Alginic acid or alginate (ALG) is a biocompatible and non-immunogenic polysaccharide obtained from kelp, brown algae, and some bacteria [14]. It is composed of two alternating blocks, α-L-guluronic acid (G) and β-D-mannuronic acid (M), linked via α- (1)(2)(3)(4) and β-(1-4) glycosidic bond, respectively ( Figure 1). Methods have evolved to obtain high purification grade ALG at a low cost, with negligible traces of contaminants, such as polyphenols and endotoxins [92,93]. Stable hydrogels can be developed in mild conditions by adding divalent metal cations (Ca 2+ , Sr 2+ , and Ba 2+ ) to aqueous ALG solution [94,95]. Its sol-gel transition is ascribed to the formation of an "egg-box" structure upon selective binding of cations to G-blocks; a phenomenon which explains the higher elastic modulus for ALG gels richer in G blocks [96].
Despite these merits, ALG is not a preferred biomaterial as it lacks cell binding motif and, therefore, exhibits poor cell adherence [97]. This has been demonstrated through a comparison between scaffolds developed from RGD-immobilized and unmodified ALG. Immobilization of the peptide promoted cell adherence to the matrix, prevented cell apoptosis, and accelerated cardiac tissue regeneration. The cardiomyocytes reorganized their myofibrils and reconstructed myofibers within six days ( Figure 2). These effects were well reflected in the expression levels of α-actinin, N-cadherin and connexin-43 in cells cultured within RGD-seeded scaffolds [98]. Enhanced cell adherence upon the attachment of RGD is explained as follows. Cellular integrins link the intracellular skeleton with ECM via the RGD peptide. It initiates the cascade for cell survival and proliferation [31]. A similar argument is applicable tothe incorporation of bone-forming peptides (derived from bone morphogenetic protein-7) into scaffolds for driving osteogenesis and osteo-differentiation [99,100].
Incorporation of poly ε-caprolactone (PCL) [101,102], CHI [103], halloysite nanotubes [104], and carbon nanotubes (CNTs) [19] has been investigated to tune the mechanical properties, bioactivity, and proliferation rate of surface cells. Herein, the specific blending ratio of components eliminates the possibility of phase separation. Acellular macroporous ALG scaffolds have shown to promote the stabilization of hepatocytes, both in vitro [105,106] and in vivo [107]. Shteyer et al. [107] demonstrated that ALG scaffolds, without implanted cells, significantly improved the survival rate of partially hepatectomized mice (87%). The animal manifested normal and prolonged aspartate-and alanine aminotransferase serum levels as compared to 2-to 20-fold increase in control groups (non-treated and collagen-treated mice). The authors correlated these findings to the non-adhesive and macroporous structure of the ALG matrix. Macroporosity enabled rapid confinement of cells within the remnant liver and caused a pronounced increase of cell polarity. Together with complimentary secretion of ECM components, growth factors, and chemokines, it created a specialized niche favorable to differentiation of remnant cells as functional hepatocytes [107]. The formation of scaffolds is dependent on pH, ion concentration, and ALG composition. Destruction of the gel network and un-controlled degradation may occur in biological buffers containing chelators or monovalent electrolytes [104,108,109]. However, scaffolds developed from covalently cross-linked ALG have shown a shape memory effect, an exploitable property while contemplating the repair of damaged annulus fibrous tissues. The formulation supported cell penetration, proliferation, and ECM deposition when cultured in intervertebral disc-like niche (low oxygen and glucose level) [110].
Alginic acid or alginate (ALG) is a biocompatible and non-immunogenic polysaccharide obtained from kelp, brown algae, and some bacteria [14]. It is composed of two alternating blocks, α-L-guluronic acid (G) and β-D-mannuronic acid (M), linked via α- (1)(2)(3)(4) and β-(1-4) glycosidic bond, respectively ( Figure 1). Methods have evolved to obtain high purification grade ALG at a low cost, with negligible traces of contaminants, such as polyphenols and endotoxins [92,93]. Stable hydrogels can be developed in mild conditions by adding divalent metal cations (Ca 2+ , Sr 2+ , and Ba 2+ ) to aqueous ALG solution [94,95]. Its sol-gel transition is ascribed to the formation of an "egg-box" structure upon selective binding of cations to G-blocks; a phenomenon which explains the higher elastic modulus for ALG gels richer in G blocks [96].
Despite these merits, ALG is not a preferred biomaterial as it lacks cell binding motif and, therefore, exhibits poor cell adherence [97]. This has been demonstrated through a comparison between scaffolds developed from RGD-immobilized and unmodified ALG. Immobilization of the peptide promoted cell adherence to the matrix, prevented cell apoptosis, and accelerated cardiac tissue regeneration. The cardiomyocytes reorganized their myofibrils and reconstructed myofibers within six days ( Figure 2). These effects were well reflected in the expression levels of α-actinin, Ncadherin and connexin-43 in cells cultured within RGD-seeded scaffolds [98]. Enhanced cell adherence upon the attachment of RGD is explained as follows. Cellular integrins link the intracellular skeleton with ECM via the RGD peptide. It initiates the cascade for cell survival and proliferation [31]. A similar argument is applicable tothe incorporation of bone-forming peptides (derived from bone morphogenetic protein-7) into scaffolds for driving osteogenesis and osteodifferentiation [99,100]. were immune stained for a-actinin (green) and nuclei (red-propidium iodide). Adjacent cardiomyocytes joined to form striated myofibers (Figure 2A, day 6), an occurrence that increased in frequency as cultivation proceeded ( Figure 2B, day 12). In contrast, cardiomyocytes cultivated within the unmodified ALG scaffolds revealed unorganized myofibrils; there were fewer interactions between adjacent cardiomyocytes and myofibers were not detected ( Figure 2C and D, days 6 and 12, The constructs were immune stained for a-actinin (green) and nuclei (red-propidium iodide). Adjacent cardiomyocytes joined to form striated myofibers (Figure 2A, day 6), an occurrence that increased in frequency as cultivation proceeded ( Figure 2B, day 12). In contrast, cardiomyocytes cultivated within the unmodified ALG scaffolds revealed unorganized myofibrils; there were fewer interactions between adjacent cardiomyocytes and myofibers were not detected ( Figure 2C and D, days 6 and 12, respectively). The lower panel shows relative locations of cardiomyocytes and nonmyocyte cells (NMCs) in (E) RGD-immobilized and (F) unmodified ALG scaffold; (G) the native adult cardiac tissue. In E and F, only cardiomyocytes were stained for α-actinin (green), while all cell nuclei were stained with propidium iodide (Red). Arrow heads denote cell nuclei of NMCs. Twelve-day constructs were fixed, fluorescently stained, and examined using confocal microscopy. In G, native adult cardiac tissue was stained for troponin-T (brown). The NMCs surrounding cardiomyocyte bundles were negatively stained. Adult rat ventricles were paraffin-fixed, cross-sectioned, and immunostained for troponin-T. Reproduced and modified with permission from Elsevier (2011) [98].

Dextran
Dextran (DEX) is a bacterially-derived uncharged, linear polysaccharide composed of α-1,6 linked D-glucopyranose residues with a few percent of α-1,2-, α-1,3-, or α-1,4-linked side chains [49]. It is available in a wide range of molecular weights and undergoes enzymatic degradation in the spleen, liver, and colon [111]. Crosslinked DEX hydrogel beads have been used for aslong as low protein-binding matrices in column chromatography [112] and in microcarrier cell culture technology [113,114]. Soft tissue-engineering applications of DEX stem from its resistance to protein adsorption and cell adhesion [115]. Porous DEX hydrogels can be prepared through crosslinking mediated by hydroxyl groups present on α-1,6-linked D-glucose residues [116]. The polymer has three hydroxyl groups in each repeat unit, and the reactivity of these groups follows the order of C 2 >C 4 >C 3 [117]. Several chemical modifications have been explored, yielding DEX derivatives with tailored physicochemical and functional characteristics [118][119][120] (Figure 3). Oxidized DEX can be crosslinked through the attachment mono-, bi-, and multiarmed amines [111,[121][122][123]. Alternatively, glycidyl methacrylate (GMA) can be attached to oxidized DEX and the latter can be crosslinked with dithiothreitol (DTT) [120].
Some groups have investigated surface grafting and co-polymerization as a tool of improving the cell-adhesion of DEX [24,124,125]. Levesque et al. [125] developed scaffolds of methacrylated-DEX copolymerized with aminoethyl methacrylate. Herein, primary amine groups served as handles to covalently link RGD peptide. The adhesion and neurite outgrowth of primary embryonic chick dorsal root ganglia increased upon copolymerization. A further improvement was noticed upon peptide immobilization. Notably, direct coupling between peptide (amine) and hemiacetals of oxidized DEX destructed the conformation of peptides. At the same time, the presence of amine-C h it o s a n Figure 3. Scheme showing the crosslinking approaches for oxidized DEX. DEX can be oxidized via periodate treatment. Oxidized DEX can be crosslinked through the attachment mono-, bi-, and multi-armed amines [111,[121][122][123]. Alternatively, glycidyl methacrylate (GMA) can be attached to oxidized DEX and the latter can be crosslinked with dithiothreitol (DTT) [120]. Some groups have investigated surface grafting and co-polymerization as a tool of improving the cell-adhesion of DEX [24,124,125]. Levesque et al. [125] developed scaffolds of methacrylated-DEX copolymerized with aminoethyl methacrylate. Herein, primary amine groups served as handles to covalently link RGD peptide. The adhesion and neurite outgrowth of primary embryonic chick dorsal root ganglia increased upon copolymerization. A further improvement was noticed upon peptide immobilization. Notably, direct coupling between peptide (amine) and hemiacetals of oxidized DEX destructed the conformation of peptides. At the same time, the presence of amine-pendants in the side-chain of constituent amino acids impaired the scaffold-cell interactions [124]. This has been minimized through the development of sulfhydryl-terminated peptides [125].
A recent study of Noel et al. [126] questioned the cell-selective response of ECM peptides using DEX scaffolds. Investigators illustrated the role of four ECM peptides (RGD, YIGSR, REDV, and CAG) upon adhesion and proliferation of HUVEC and AoSMC cells. A library of vinylsulfone-modified DEX was tethered with the peptides. RGD (Arg-Gly-Asp), YIGSR (Tyr-Ile-Gly-Ser-Arg), and SGIYR (Gly-Ile-Tyr-Arg) were able to enhance both HUVEC and AoSMC adhesion (showing no selectively for HUVEC over AoSMC), whereas REDV (Arg-Glu-Asp-Val) and CAG (Cys-Ala-Gly) failed in improving the cell adhesion. Interestingly, co-immobilization of vascular endothelial growth factor and RGD resulted in selective proliferation of HUVEC cells. It thus highlighted the scope of changing the conformation, sequence tuning, and lengthening of peptides as tactics to impart a cell selective response in the scaffold.

Hyaluronic Acid
Commercial hyaluronic acid (HA) is extracted from rooster combs, but it has also been produced using genetically engineered bacteria. Highly pure HA is available in a range of molecular weights at relatively low costs. HA and its derivatives are widely used in the cosmetics industry, medicine, and surgery. Physiochemical and biological properties, methods of modification, and drug delivery applications of HA have been described in other comprehensive reviews [52,127,128]. Its biological activity is molecular weight-dependent [129]; high molecular weight HA has anti-inflammatory and anti-angiogenic properties, whereas low-molecular weight HA possesses pro-inflammatory and pro-angiogenic activities [130][131][132]. Besides, studies show that HA promotes macrophage differentiation into the M2 phenotype [133]. Improved cellular proliferation and tissue regeneration have been demonstrated by blending with biodegradable materials [134][135][136] and coating the scaffolds with HA [137,138] and non-covalent binding [139]. These events are most likely mediated through selective interaction of HA with cell surface receptors, such as CD44, ICAM-1, and RAHMM [52,140].
Kudryavtseva et al. [141] explored the effect of surface immobilized high molecular weight HA upon survival of primary human monocyte-derived macrophages. The immobilization on polylactic acid scaffolds was accomplished through atmospheric pressure cold plasma treatment. HA attachment enhanced the biocompatibility of the scaffold and stimulated its pro-angiogenic action. Interestingly, dip coating of HA (1 wt% solution) has been shown to enrich the attachment of MCF7 cells onto poly(lactic acid-co-glycolic acid) (PLGA) scaffolds [137]. Depending on the process parameters, deposited HA may have configurations ranging from thin disconnected aggregates to a thick continuous layer on the pore surface ( Figure 4). Besides, layer topography may affect the swelling of scaffold and may be of interest in applications wherein resistance to normal stress is desirable [138]. For other specific applications, hybrid nanofibres can be used as reinforcement alternative [142].
While the majority of investigations have focused on exploiting the direct biological effects of HA, its incorporation intriguingly improved the mechanical strength of scaffolds and may, therefore, inhibit the cell-induced contractions. Davidenko et al. [143] investigated the influence of increasing the amount of HA upon mechanical characteristics of collagen scaffolds. Together with supporting the proliferation of 3T3-L1 preadipocytes, HA created additional crosslinks. Consequently, the scaffold exhibited improved resistance to compression and in vitro dissolution.  C). These aggregates become more distinct as the number of cycles increases, but a uniform layer is not obtained with 0.5wt% even after five cycles (D). In contrast, coating with 5wt% HA produces a uniform continuous layer after the first coating cycle. The effect of further cycles is to achieve the layer thickness. This is accompanied with a decreased pore diameter and the clogging of some pores (E). After the fifth cycle (F), the channels are filled with HA to a high degree. (05HA# and 5HA# designate the scaffolds coated, respectively, with 0.5wt% and 5wt% HA solutions, # being the number of cycles). Reproduced and modified with permission from Elsevier (2011) [138].
While the majority of investigations have focused on exploiting the direct biological effects of HA, its incorporation intriguingly improved the mechanical strength of scaffolds and may, therefore, inhibit the cell-induced contractions. Davidenko et al. [143] investigated the influence of increasing the amount of HA upon mechanical characteristics of collagen scaffolds. Together with supporting the proliferation of 3T3-L1 preadipocytes, HA created additional crosslinks. Consequently, the scaffold exhibited improved resistance to compression and in vitro dissolution.

Approaches of Scaffold Preparation
Prototype scaffold preparations include three key components: Support material, cells, and angiogenic factors. Typically, a blend of biopolymers is employed with the objectives, such as enhancing mechanical properties, and tuning the porosity, loading property, swelling ratio, and degradation kinetics of the scaffold. Cells and growth factors either adhere to the scaffold surface [144] or get encapsulated within the matrix [145]. Formulations include hydrogel [5,91], fiber [142], film [69], and de-cellularized matrices [146][147][148]. Electrospun microfiber bundles are suturable and  C). These aggregates become more distinct as the number of cycles increases, but a uniform layer is not obtained with 0.5wt% even after five cycles (D). In contrast, coating with 5wt% HA produces a uniform continuous layer after the first coating cycle. The effect of further cycles is to achieve the layer thickness. This is accompanied with a decreased pore diameter and the clogging of some pores (E). After the fifth cycle (F), the channels are filled with HA to a high degree. (05HA# and 5HA# designate the scaffolds coated, respectively, with 0.5wt% and 5wt% HA solutions, # being the number of cycles). Reproduced and modified with permission from Elsevier (2011) [138].

Approaches of Scaffold Preparation
Prototype scaffold preparations include three key components: Support material, cells, and angiogenic factors. Typically, a blend of biopolymers is employed with the objectives, such as enhancing mechanical properties, and tuning the porosity, loading property, swelling ratio, and degradation kinetics of the scaffold. Cells and growth factors either adhere to the scaffold surface [144] or get encapsulated within the matrix [145]. Formulations include hydrogel [5,91], fiber [142], film [69], and de-cellularized matrices [146][147][148]. Electrospun microfiber bundles are suturable and often exhibit an elastic modulus identical to that of native tissue [9]. Transplantation can be rendered less aggressive by developing in situ gelling formulations, which later acquire the configurations of damaged tissue [23,149]. Besides, self-crosslinking has been achieved in neat polysaccharide systems via thiolation [25].
The preparation method must be selected on the criteria, such as a desired scale of operation, controllability of steps, and batch-to-batch consistency. A general approach includes dissolving the component(s) into an aqueous vehicle and subsequent processing via solvent casting, lyophilization, electrospinning, or cryo-gelation. The weight ratio of components is adjusted to attain a desired dispersibility [57]. This is essential with the consideration that cross-linkage between the constituents may sometimes offset the hydrophilicity and pore size of polysaccharide scaffolds [74,150].
The scaffold can be macro-or micro-patterned at a high accuracy using controlled chemical manipulations to achieve desirable biophysical characteristics [151]. Photo-crosslinkable interpenetrating (IPNs) and semi-interpenetrating networks (SIPNs) between COL and HA have been shown to control the structural and biomechanical properties ( Figure 5I). In contrast to IPN composed of two un-crosslinked polymers exhibiting full interpenetration, SIPN consists of one crosslinked polymer entangled in another un-crosslinked polymer and hence, is mechanically inferior [52]. Such entangled networks retain the structural properties of component polymers while reinforcing the scaffold. Scaffolds developed from the IPN-SIPN blend are anisotropic; showing region-specific distribution of crosslinking density, viscoelasticity, water content, and porosity [151,152] (Figure 5II). often exhibit an elastic modulus identical to that of native tissue [9]. Transplantation can be rendered less aggressive by developing in situ gelling formulations, which later acquire the configurations of damaged tissue [23,149]. Besides, self-crosslinking has been achieved in neat polysaccharide systems via thiolation [25]. The preparation method must be selected on the criteria, such as a desired scale of operation, controllability of steps, and batch-to-batch consistency. A general approach includes dissolving the component(s) into an aqueous vehicle and subsequent processing via solvent casting, lyophilization, electrospinning, or cryo-gelation. The weight ratio of components is adjusted to attain a desired dispersibility [57]. This is essential with the consideration that cross-linkage between the constituents may sometimes offset the hydrophilicity and pore size of polysaccharide scaffolds [74,150]. Figure 5. Schematic for synthesizing IPNs, SIPNs, and photopatterned hydrogels (I). HA and collagen solution were suspended in the silicone mold and collagen was permitted to undergo fibrillogenesis at 37°C (1). This resulted in the formation of a SIPN, which was then exposed to ultra-violet light to yield a full IPN (2). Alternatively, photo patterning was performed using a photomask, which resulted in SIPN and IPN patterns within a single hydrogel (3) (I). The lower panel (II) shows the macro-and micro-patterned hydrogels formed due to differential crosslinking densities. A macropatterned hydrogel is shown in which half was exposed to UV before (A) and after washing the un-crosslinked fluorescein acrylate (B). In addition, (B) shows the interface between the macropatterned halves. Micropatterning within a single bulk hydrogel of a 500 μm thickness is shown in C and D. (Scale bar-150 μm). Reproduced and modified with permission from Elsevier (2009) [151].
The scaffold can be macro-or micro-patterned at a high accuracy using controlled chemical manipulations to achieve desirable biophysical characteristics [151]. Photo-crosslinkable I II Figure 5. Schematic for synthesizing IPNs, SIPNs, and photopatterned hydrogels (I). HA and collagen solution were suspended in the silicone mold and collagen was permitted to undergo fibrillogenesis at 37 • C (1). This resulted in the formation of a SIPN, which was then exposed to ultra-violet light to yield a full IPN (2). Alternatively, photo patterning was performed using a photomask, which resulted in SIPN and IPN patterns within a single hydrogel (3) (I). The lower panel (II) shows the macroand micro-patterned hydrogels formed due to differential crosslinking densities. A macropatterned hydrogel is shown in which half was exposed to UV before (A) and after washing the un-crosslinked fluorescein acrylate (B). In addition, (B) shows the interface between the macropatterned halves. Micropatterning within a single bulk hydrogel of a 500 µm thickness is shown in C and D. (Scale bar-150 µm). Reproduced and modified with permission from Elsevier (2009) [151].
Khoshakhlaghet al. [153] illustrated the effects of micro-patterning upon neurite growth using a dual hydrogel, incorporating methacrylated HA and Puramatrix (PM, a self-assembling peptide scaffold). Initially, IPN hydrogels were formulated using self-assembly of PM and photo-crosslinking of HA. It was then surrounded by photo-crosslinkable polyethylene glycol (PEG). Integration between the two compartments of hydrogel was mediated by the IPN. Crosslinkable substrates were exposed to UV radiations in geometries relevant to cover the entire gel thickness, thereby creating a desirable micro-patterning. A range of mechanical properties could be achieved by controlling the degree of methacrylation. Regions with a lesser degree of methacrylation (greater porosity) displayed better neurite outgrowth [153] (Figure 6).
inferior [52]. Such entangled networks retain the structural properties of component polymers while reinforcing the scaffold. Scaffolds developed from the IPN-SIPN blend are anisotropic; showing region-specific distribution of crosslinking density, viscoelasticity, water content, and porosity [151,152] (Figure 5II). Khoshakhlaghet al. [153] illustrated the effects of micro-patterning upon neurite growth using a dual hydrogel, incorporating methacrylated HA and Puramatrix (PM, a self-assembling peptide scaffold). Initially, IPN hydrogels were formulated using self-assembly of PM and photo-crosslinking of HA. It was then surrounded by photo-crosslinkable polyethylene glycol (PEG). Integration between the two compartments of hydrogel was mediated by the IPN. Crosslinkable substrates were exposed to UV radiations in geometries relevant to cover the entire gel thickness, thereby creating a desirable micro-patterning. A range of mechanical properties could be achieved by controlling the degree of methacrylation. Regions with a lesser degree of methacrylation (greater porosity) displayed better neurite outgrowth [153] (Figure 6). Figure 6. A significant increase (p<0.05) in the swelling ratio can be noticed in IPN32, with a smaller crosslinking density (A). Besides, a higher crosslinking density (IPN90) led to a significant increase (p<0.05) in the compressive moduli (B). Comparative analysis of the length of neurite extension in IPN90 and IPN32 constructs is shown in (C) and (D). A less stiff substrate allowed a longer growth, with some neurites extending up to 3.3 mm after 7 days (C). Analysis of the amount of neurite growth (average of five longest neurites) demonstrated that the more compliant substrate allowed superior overall growth (D). Reproduced and modified with permission from Elsevier (2015) [153].
Techniques, such as embossing, micro-contact printing, and layer-by-layer assembly of planer sheets, have been employed for the fabrication of micro-patterned scaffolds [154,155]. Scaffolds with honeycomb, square, and rectangle patterns (needed for specialized applications) are obtainable using these methods [156,157]. For instance, it is desirable from the cardiac scaffolds to offer electrical cues, in addition to biomimicking mechanical and topographical features. Liu et al. [158] fabricated micropatterned cardiac patches using a tri-culture system, composed of cardiomyocytes in combination with cardiac fibroblasts and endothelial cells. Techniques, such as embossing, micro-contact printing, and layer-by-layer assembly of planer sheets, have been employed for the fabrication of micro-patterned scaffolds [154,155]. Scaffolds with honeycomb, square, and rectangle patterns (needed for specialized applications) are obtainable using these methods [156,157]. For instance, it is desirable from the cardiac scaffolds to offer electrical cues, in addition to biomimicking mechanical and topographical features. Liu et al. [158] fabricated micro-patterned cardiac patches using a tri-culture system, composed of cardiomyocytes in combination with cardiac fibroblasts and endothelial cells.
3D printing technology is also gaining popularity for its high speed and continuous scaffold design. Typically, a bioink containing cells, growth factors, and other biological solutions is printed over acellular scaffolds. A highly customized architecture can be achieved with the help of a motion-controlled multinozzle deposition system [19]. The process employs a low pressure extrusion and is operated at room temperature, with benign processing requirements. Low pressure extrusion with a large diameter nozzle helps in minimizing mechanical stress to the cells. It is, however, important that the material is sufficiently viscous to be dispensed as free standing filaments exhibiting desirable mechanical strength. The reader is referred to earlier reviews on the application of 3D printing technology in tissue engineering [159][160][161].

Clinical Status of Polysaccharide Scaffolds
The evaluation of scaffolds in a clinical set-up is necessary to validate its efficacy. Clinical reports on polysaccharide scaffolds are interesting, but the power of those findings is limited due to a small sample size, lack of a randomized control group for comparison, and the unavailability of long-term studies [162,163]. With a limited sample size and smaller follow-up period, the investigators may miss infrequent adverse events. In this landscape, the regenerative response with the test approach remains obscure. On the other hand, long-term data, acquired in a broader population, provide important indications if early risks associated with the intervention can be offset by future benefits [26,163].
Stillaert et al. [164] investigated HA-based preadipocyte-seeded scaffolds for adipo-conductive potential and efficacy in humans. Autologous cells, isolated from lipoaspirate material and seeded on HA scaffolds, were implanted subcutaneously. The scaffold displayed superior cellularity and progressive tissue integration within eight weeks of implantation. It, however, lacked angiogenic penetration since the cells were located more than 100 µm away from the native micro vasculature; beyond the diffusive capacity of oxygen [164]. The adherence of the scaffold to the lesion can be monitored by analyzing the polysaccharide content in biopsy samples [165].
Other clinical studies have employed esterified HA (HYAFF ® ). Esterification of carboxyl groups involved the preparation of a quaternary HA salt and its subsequent reaction with an esterifying agent (aliphatic, alicyclic, or aromatic alcohol) in an aprotic solvent [166]. Scaffolds based on benzyl ester (HYAFF ® 11) have been widely tested for cartilage repair ( Table 2). The treatment minimizes pain and counteracts the development of arthritis [167]. It is agreed that the autologous cell-based repair technique results in the generation of hyaline-like repair tissue. It shows a lower probability of failure in comparison to fibrous repair tissue [168]. This might be the reason for the greater clinical acceptability of scaffold-based cell seeding over bone-marrow stimulating techniques for cartilage repair [165,167]. regeneration occurred in about 50% of patients during 6 to 30 month follow-up [171] ALG has been clinically tested for improving cardiac function using Algisyl-LVR™ [162] and IK-5001 technologies [163]. Lee et al. [162] employed a proprietary gel, which transformed to a scaffold upon placement in the affected region. The formulation consisted of: (a) ALG component as 4.6% aqueous mannitol, and (b) Ca 2+ -ALG component as insoluble particles suspended in 4.6% aqueous mannitol. These solutions were extemporaneously mixed in one syringe prior to intramyocardial administration [162] (Figure 7). On the contrary, IK-5001 comprises of 1wt% ALG containing 0.3% calcium gluconate and undergoes in situ crosslinking. Its intracoronary delivery is relatively simple and does not require a unique device or complex imaging system. When injected, the formulation selectively permeates to the infracted myocardial tissue and reversibly crosslinks to form a temporary bioabsorbable cardiac scaffold in a Ca 2+ dependent manner. The scaffold then replaces the damaged ECM, reduces myocardial wall thinning and strain, and ultimately attenuates infarct expansion [163,172]. Herein, the selectivity of scaffold deposition is ascribed to abnormal microvascular permeability and elevated extracellular Ca 2+ concentrations within the infarct zone [173].
hyaline cartilage regeneration women), mean age-35.6 years, mean follow-up -16 months (range, 6-30); mean implant area-5.0 cm 2 regeneration occurred in about 50% of patients during 6 to 30 month follow-up [171] ALG has been clinically tested for improving cardiac function using Algisyl-LVR™ [162] and IK-5001 technologies [163]. Lee et al. [162] employed a proprietary gel, which transformed to a scaffold upon placement in the affected region. The formulation consisted of: (a) ALG component as 4.6% aqueous mannitol, and (b) Ca 2+ -ALG component as insoluble particles suspended in 4.6% aqueous mannitol. These solutions were extemporaneously mixed in one syringe prior to intramyocardial administration [162] (Figure 7). On the contrary, IK-5001 comprises of 1wt% ALG containing 0.3% calcium gluconate and undergoes in situ crosslinking. Its intracoronary delivery is relatively simple and does not require a unique device or complex imaging system. When injected, the formulation selectively permeates to the infracted myocardial tissue and reversibly crosslinks to form a temporary bioabsorbable cardiac scaffold in a Ca 2+ dependent manner. The scaffold then replaces the damaged ECM, reduces myocardial wall thinning and strain, and ultimately attenuates infarct expansion [163,172]. Herein, the selectivity of scaffold deposition is ascribed to abnormal microvascular permeability and elevated extracellular Ca 2+ concentrations within the infarct zone [173]. Altogether, clinical applications of polymeric scaffolds are challenging due to intricacies of replicating the complex tissue environment without eliciting undesirable immunologic events. Tissue remodeling is often constrained by limited diffusion of oxygen and growth factors in polymeric scaffolds [174]. This has been demonstrated with reference to vascular restorative therapy in cardiac tissues. Herein, the bioresorbable nature of scaffold stimulated positive blood vessel wall remodeling and restoration of contractile functions. With these considerations, polymeric scaffolds appear Altogether, clinical applications of polymeric scaffolds are challenging due to intricacies of replicating the complex tissue environment without eliciting undesirable immunologic events. Tissue remodeling is often constrained by limited diffusion of oxygen and growth factors in polymeric scaffolds [174]. This has been demonstrated with reference to vascular restorative therapy in cardiac tissues. Herein, the bioresorbable nature of scaffold stimulated positive blood vessel wall remodeling and restoration of contractile functions. With these considerations, polymeric scaffolds appear superior to metallic stents. The latter often lead to distorted vessel physiology, incomplete endothelialization, and stent fracture [175,176]. At the same time, we cannot overlook the fact that behavior of polymeric materials may not be identical under dry and submerged conditions. It is, therefore, suggestive to preliminarily map the localized changes in structural integrity vis-a-vis macroscopic performance of scaffolds under practical use conditions [177]. This can certainly minimize the clinical failure of polymeric scaffolds.

Conclusions and Perspective
Preservation of local organ function via tissue regeneration is a definitive component of post-operative care. Regenerative treatment leans over the development of advanced biomaterials, processing thereof as 3D scaffolds, and investigating the manner in which scaffold material cooperates with the seeded cells, proteins, and growth factors in order to augment the natural repair mechanisms.
Polysaccharide based materials have shown endless promise for developing the tissue scaffolds. Ongoing advancements in polysaccharide chemistry and nanotechnology have enabled the integration of mechanical, topographical, and biological cues into these materials for stably recapitulating the tissue-scale organization. Importantly, these recreated structures may act as customized tissue surrogates for the screening of new drug molecules [178].
While debating the translational application of scaffolds, it is of interest to look forward to upcoming technologies for mass production and setting quality control parameters. For instance, studies illustrate that a polymer's molecular weight and sterilization procedure affect biological and microstructural attributes of the scaffold [179,180]. Therefore, in sync with the advances in polymer modification approaches, investigations must equally focus on the way these modifications affect the architecture and biological performance. This would enable the development of preparations compatible with regulatory standards worldwide and those showing a lesser rejection in clinical settings. Results on larger patient cohorts will indeed show the footprints of scaffold research on clinical medicine in the future.
Author Contributions: P.B. conceived and designed the structure of this review. R.P. contributed to literature survey. S.T. wrote the manuscript which was finally edited by P.B.