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Tissue engineering aims to regenerate tissues and organs. Many tissue engineering approaches involve three dimensional (3D) scaffolds [39]. A scaffold provides a hospitable environment for tissue regeneration. It should be biocompatible to allow elaborate multi-cellular processes to be carried out, besides being in concert with cell and tissue-specific events such as cell proliferation, migration, and differentiation. Furthermore, degradation products of a scaffold should not induce local or systemic adverse events to avoid future complications. In addition to stiffness and elasticity, an ideal scaffold is also expected to show desired porosity and pore interconnectivity properties to enable the metabolic exchange, waste disposal, colonization, and survival of entrapped cells [40]. Natural polysaccharides such as alginate [41], collagen [39], chondroitin sulfate [42], chitosan [43], and hyaluronic acid [44] have been used to generate scaffolds. Since scaffolding materials protect their contents from the surrounding biological environment, scaffolds are also used for the delivery of therapeutics, growth factors, and even therapeutically useful cells [45].

In recent years, the production of biocompatible scaffolds composed of decellularized cellulose combined with hydrogels and biopolymers has gained attention in the field of biomaterial science [46, 47]. Cellulose is abundant in nature and it can be obtained and produced easily. Cellulose is known for its biocompatibility, bioactivity, sustainability, and eco-efficiency properties. Thereby, to minimize the utilization of animal and human-derived biomolecules, cellulose-based materials have great potential to become the next generation of green chemistry-based biomaterials as an alternative to conventional polymers [48, 49]. However, it should be pointed out that cellulose appears as a very slowly degradable even non-degradable material. Märtson et al. reported that degradation of viscose cellulose sponges implanted subcutaneously into rats took longer than 60 weeks because of the absence of enzymes that attack the β (1→4) linkage [50]. However, an ideal scaffold should be constructed from materials degradable in the organism for replacement by natural extracellular matrix. Oxidation of cellulose (i.e., achieved by using various oxidizing agents, such as nitrogen oxides, free nitroxyl radicals, NaClO₂, or CCl₄) is one of the methods to increase the degradability since the oxidized polymer readily undergoes chain shortening to give oligomers which will be further hydrolyzed to smaller fragments, including glucuronic acid and glucose by hydrolytic enzymes [51, 52]. The oxidation of cellulose converts glucose residues to glucuronic acid residues containing –COOH groups which modulate the degradation kinetics of cellulose, its pH, its swelling capacity in a water solution, and mechanical stability. Additionally, the polar and negatively charged nature of the –COOH groups facilitate the oxidized cellulose to be used for functionalizing with various biomolecules [51].

Although cellulose is mainly obtained from vegetal products, it is also produced extracellularly by bacteria from the genera Gluconacetobacter, Sarcina, and Agrobacterium [53]. The most prominent and well-known bacterial cellulose producing species is Komagataeibacter xylinus (Gluconacetobacter xylinus). K. xylinus is not the sole species among acetic acid bacteria with tremendous potential for bacterial cellulose production, since also other species, like Komagataeibacter medellinensis, Komagataeibacter hansenii, Komagataeibacter oboediens, Komagataeibacter nataicola, Komagataeibacter rhaeticus, Komagataeibacter pomaceti, and Komagataeibacter saccharivorans have been shown as stupendous cellulose producers. Of note, acetic acid bacteria species used for cellulose production are considered as “generally recognized as safe-GRAS” [54, 55]. Previous studies have shown that through oxidative fermentation, cellulose producing bacteria can use various types of sugars in both synthetic and non-synthetic media as carbon sources [54, 56]. Cellulose is formed as a thick gel or pellicle on the surface of the growth medium in stationary culture conditions. This cellulose is different from plant cellulose in its degree of polymerization, crystallinity, higher purity, and tensile strength. Besides, unlike the vegetal cellulose that presents mainly the Iβ structure, bacterial cellulose has Iα and Iβ crystalline forms. Biocompatibility property of bacterial cellulose made it suitable for several biomedical applications [57]. Similar to plant cellulose, modifications of bacterial cellulose has been investigated to utilize it for biomedical applications. These efforts include processes for changing the surface chemistry of bacterial cellulose by incorporating different substrates such as small molecules, inorganic nanoparticles, and polymers to improve its functionality [57, 58]. In addition, to be used for tissue engineering, several composites were designed by taking the advantage of porous structure and mechanical strength of bacterial cellulose, such as composites of bacterial cellulose/collagen, bacterial cellulose/agarose, bacterial cellulose/poly(3-hydroxybutyrate) (PHB), bacterial cellulose/chitosan and bacterial cellulose/hydroxyapatite (Hap) [54]. However, it should be mentioned that like plant cellulose, the high crystalline nature and absence of enzyme which could break β (1→4) glycosidic linkage of bacterial cellulose in the human body prevent its degradation in vivo [54]. Although this feature of cellulose is advantageous for providing long term support as a scaffold material, a scaffold should degrade in time to allow the healing and regeneration process. Besides this, degradation products should be biocompatible [59]. Different approaches have been evaluated to enhance the degradability of bacterial cellulose such as chemical methods including modification of bacterial cellulose by periodate oxidation [60]. In another study, aldehyde groups were introduced to bacterial cellulose nanofibers [61]. Though, Yadav et al. used a metabolic engineering-based approach to induce the degradability of cellulose: N-acetylglucosamine (GlcNAc) residues were introduced into cellulosic biopolymers during de novo synthesis from Gluconacetobacter xylinus. The presence of GlcNAc enabled bacterial cellulose to be susceptible to lysozyme and also disrupted the highly ordered cellulose crystalline structure. Accordingly, in vivo studies showed that modified cellulose from the engineered strain was almost entirely degraded at day 10 and was completely undetectable at day 20 while little to no degradation of the cellulose obtained from the control bacteria at either time point [62].

Gelatin, a denatured polypeptide product of collagen, is a promising biomaterial as a scaffold. It has some unique regenerative characteristics, including its chemical similarities to the native extracellular matrix, low antigenicity, biocompatibility, and biodegradability. Besides, gelatin is cheap and abundant, and also has accessible functional groups that allow chemical modifications with other biomolecules [63]. However, low solubility in concentrated aqueous media, high viscosity, poor mechanical properties, and sensitivity to enzymatic degradation properties limit its applications as a scaffold material [63, 64]. The approach of combining gelatin with a wide range of polysaccharides has been used to overcome these drawbacks of gelatin. Hydrogels are three-dimensional cross-linked porous networks that consist of biopolymers or polyelectrolytes and would swell with a large amount of water or biological fluid [65]. Hybrid hydrogels composed of gelatin and polysaccharides, mainly cellulose, alginate, and hyaluronic acid, provides green and natural platforms for cell and tissue engineering. The combination of gelatin with polysaccharides is advantageous to better mimic the proteoglycan containing extracellular matrix. In addition, gelatin-polysaccharide biomaterials have been described to show biocompatibility, mechanical resilience, high stability, low thermal expansion, antimicrobial and anti-inflammatory properties [63].

Hyaluronic acid, also called hyaluronan, is an acidic, non-sulfated glycosaminoglycan present throughout the human body. Hyaluronic acid maintains the viscoelasticity of the extracellular matrix, therefore supports cellular structure and functions. It also keeps tissues hydrated and maintains the integrity of the extracellular matrix. Mechanistically, hyaluronic acid is known to interact with the receptors CD44, Intercellular Adhesion Molecule 1 (ICAM-1), and Hyaluronan-mediated motility receptor (HMMR), and these receptor-ligand interactions have been shown to regulate cell behaviors such as motility and adhesion [40, 66]. Hyaluronic acid is receiving special attention in a broad range of applications including cosmetics industry, biomedical and tissue engineering applications. As a main component of the extracellular matrix, hyaluronic acid is involved in tissue repair and displays advantageous physical–chemical properties, like biodegradability, biocompatibility, and viscoelasticity. Commercially, hyaluronic acid has been isolated from rooster combs; besides, it has also been produced using genetically modified bacteria [40, 67]. Biological activity of hyaluronic acid depends on its molecular weight: high molecular weight hyaluronic acid has been evaluated to show a pro-resolving response, while low molecular weight hyaluronic acid is known with its pro-inflammatory and pro-angiogenic activities [68]. It has been hypothesized that molecular weight dependent physiological effects of hyaluronic acid can be caused by an interaction between hyaluronic acid and certain receptors via different states of aggregation [69]. Nevertheless, hyaluronic acid has some disadvantages including short turnover and poor mechanical properties. Therefore, chemical modification or crosslinking approaches targeting carboxyl groups, hydroxyl group, and –NHCOCH₃ of hyaluronic acid have been studied to overcome these limitations [70]. Investigation and manufacturing composite scaffolds to improve cell viability, proliferation, attachment, differentiation, vascularization, and host integration properties have been gaining attention [66, 67]. For example, a biomimetic scaffold consisting of a bioglass–collagen–hyaluronic acid–phosphatidylserine composite has been evaluated to enhance the adhesion, proliferation, and migration properties of human mesenchymal stem cells. In another study, hyaluronic acid, silk fibroin, and collagen combinations showed to be osteogenetic [66]. Hyaluronic acid-based materials are also used in hydrogel form to obtain high water content, oxygen, nutrients, and metabolites permeable scaffolds. For instance, Zanchetta et al. designed a hydrogel scaffold based on hyaluronic acid, chondroitin 6 sulfate, and dermatan sulfate with a promising osteogenesis-promoting property in rat models [71]. Hyaluronic acid is also considered as a promising candidate for central neural tissue engineering, because of its interconnected porous structure which facilitates the delivery of nutrition and penetration of cells, nerve fibers and blood vessels. In in vivo models, hyaluronic acid was demonstrated to be effective in reducing glial and peripheral scar formation and enhancing neural regeneration [72, 73]. The modulus of hyaluronic acid hydrogels was also reported to affect differentiation of neural progenitor cells: most of the neural progenitor cells cultured in hydrogels with mechanical properties comparable to those of neonatal brain tissue differentiated into neurons with extended long, branched processes, however neural progenitor cultured in stiffer hydrogels, with mechanical properties comparable to those of adult brain, mostly differentiated into astrocytes [74].

Alginate, also known as alginic acid, is a non-immunogenic, biocompatible polysaccharide obtained from kelp, brown algae, and some bacteria. It is possible to obtain stable hydrogels of alginate through the addition of divalent metal cations, such as Sr²⁺, Ca²⁺, and Ba²⁺, to aqueous alginate solution [40, 75]. However, pure alginate has been shown to have several drawbacks including poor efficiency for promoting cell adhesion and slow degradation kinetics in physiological conditions [76, 77]. Therefore, studies have been focused on enhancing its physicochemical properties and cytocompatibility. For instance, Sarker et al. developed an alginate-gelatin crosslinked hydrogel through covalent crosslinking of alginate di-aldehyde with gelatin and showed that this composite hydrogel supported cell attachment, spreading and proliferation of human dermal fibroblasts [76]. In another study, immobilization of the arginylglycylaspartic acid (RGD) peptide, the most common peptide motif responsible for cell adhesion to the extracellular matrix, to sodium alginate was found to promote cell adherence to the matrix, accelerate cardiac tissue regeneration while preventing cell apoptosis [78]. On the other hand, poor solubility and slow degradation rate of pure alginate are other limitations in the use of it in biomedical applications. Although mammals do not have enzymes to degrade alginate, via exchange reactions involving monovalent cations such as sodium ions, ionically cross-linked alginate gels can be dissolved by the liberation of the divalent ions cross-linking the gel into the surrounding media. However even if the gel dissolves, it may not be cleared from the body [79]. An attractive approach to increase the biodegradability of alginate includes oxidation of alginate chains, typically with sodium periodate without interfering with its gel-forming ability in the presence of divalent cations [80].

Dextran is an uncharged, linear homopolysaccharide synthesized from sucrose by several lactic acid bacteria including Streptococcus mutans, Leuconostoc mesenteroides, and Lactobacillus brevis. High water solubility, biocompatibility, biodegradability, non-immunogenicity, and non-antigenicity properties make dextran as a good candidate for biomedical applications, including tissue engineering and drug delivery. First, dextran can be biodegraded easily by the enzyme dextranase presents in several organs of the human body including liver, colon, spleen, and kidney. Dextran can also be metabolized by different bacteria residing in the human colon [81]. To be used in biomedical applications, native dextrans with high molecular weights hydrolyzed by partial depolymerization [82]. Porous dextran hydrogels have been produced through crosslinking reactions mediated by hydroxyl groups of α-1,6-linked d-glucose residues and numerous other chemical modifications yielding dextran derivatives have been also explored. Furthermore, researchers have used co-polymerization and surface grafting methods to improve the cell-adhesion property of dextran [40]. Noel et al. investigated the cell-selective response of extracellular peptides using dextran scaffolds and they showed that vinylsulfone-modified dextran tethered with the peptides RGD (Arg-Gly-Asp), YIGSR (Tyr-Ile-Gly-Ser-Arg), and SGIYR (Gly-Ile-Tyr-Arg) was able to improve cellular adhesion [83]. In another study Sun et al. modified dextran hydrogel by decreasing crosslinking density and therefore they improved the hydrogel properties including increased swelling, rapid disintegration, reduced rigidity, and increased vascular endothelial growth factor release capability. In addition, immobilization of defined angiogenic growth factors in that modified dextran macroporous scaffold was capable to induce a rapid proliferation of functional vasculature, in vivo [84].

Gellan gum is another bacterial polysaccharide used in biomedical applications. During fermentation, Sphingomonas strains including Sphingomonas elodea (ATCC31461), Sphingomonas paucimobilis NK2000, Sphingomonas paucimobilis E2 (DSM 6314), and Sphingomonas paucimobilis GS1 can produce this high molecular weight, anionic, and linear extracellular polysaccharide. Based on the number of acetyl groups, gellan gum can be classified as native gellan gum, deacetylated gellan gum and, deacetylated and clarified gellan gum [81, 85]. The native form of gellan is composed of acetyl and L-glyceryl groups bounded to glucose residue adjacent to glucuronic acid. In the fermentation broth, acetyl and L-glyceryl groups can be eliminated by hot alkaline hydrolysis to obtain a deacylated, linear, simple polysaccharide chain generically known as gellan gum. This deacetylation process leads to a change in the gellan form from a soft, flexible, thermally reversible structure to a more rigid, more brittle and more thermo-resistant structure. Whereas, to obtain clarified gellan gum, deacylated gellan is subjected to a clarification process by increasing the temperature of the fermentation broth to 95 °C. This heating process causes the killing of bacterial cells, getting rid of cell protein residues, and reducing the viscosity of the broth. The process is completed by filtration followed by alcohol precipitation [81]. Depending on the purity levels, there are different trades of gellan gum to be used in pharmaceutical applications, food or pharmaceutical industries or preparation of biological growth media for plant tissue cultures and microbial cultures [86]. With its high water-retaining, viscoelasticity, high biocompatibility, and biodegradability properties, gellan hydrogels can be used as scaffold materials. When injected into the defective tissue area, a hydrogel of gellan can suit the shape of the defect in its gel structure. This feature of gellan hydrogels makes them ideal to be used in regenerative medicine. Accordingly, gellan-based hydrogels have been extensively examined in the context of tissue engineering applications including disc regeneration, tendon repair, fibrocartilage tissue engineering, spinal cord repair, neoskin vascularization, and wound healing along with osteochondral and bone tissue regeneration [87–91]. However, the mechanical properties and processability of gellan gum are not satisfactory for tissue engineering: gellan gum hydrogels are mechanically weak, their high gelling temperature is unfavorable and specific attachment sites for anchorage-dependent cells are inadequate. However, chemical modification and functionalization through the incorporation of the multiple hydroxyl groups and the free carboxyl per repeating unit of gellan gum has been claimed to be used to optimize its physicochemical and biological properties [92].

Recently, bioprinting has emerged as a potentially revolutionizing method for personalized regenerative medicine [93]. Using the bioprinting method, it is possible to generate 3D tissues and organs as direct copies of patienťs organ parameters [94]. A bioink has to be printable, but also should provide the required elasticity and strength to mimick the mechanical properties of native tissues and maintain the original printed structure for a long time. Besides, a bioink has to be biocompatible and should not cause an inflammatory reaction, and it has to support attachment, proliferation, and differentiation of cells. Sometimes biodegradability is aimed; if it is the case, degradation products have to be evaluated for cytotoxicity [93, 95]. Natural polysaccharides, such as collagen, hyaluronic, acid, alginate, and different hydrogels, can be utilized as 3D bioprinting materials for the printing of various types of structures as scaffolds [96]. Furthermore, synthetic polymers such as polycaprolactone are used to obtain an optimal mechanical strength of the printed constructs [95]. Designing natural polysaccharide-based 3D-printed technologies/scaffolds needs detailed investigations about their cell compatibility, flexibility, and degradation. Although 3D bioprinting is a promising technique for tissue engineering and regenerative medicine, developments in the biomaterial ink field still need further improvements to be used in clinical trials.