**3. Choice of material in artificial supporting matrixes**

The involvement of suitable materials, whether synthetic or naturally extracted in the fabrication of the artificial tissue environment, remains a challenge. The prerequisite in the development of 3D constructs includes biocompatibility, biodegradability, mechanical strength, interconnected and antimicrobial porous mechanical strength, antimicrobial and interconnected porous networks in which cellular activities could be performed analogously to their native tissue domicile. Taking into account the fact that biomaterials that actually contain the structural

component or similar biochemical and physicochemical identity of native tissue have been granted for the processing of the platform in support of artificial cells. Meanwhile, the involvement of any unique material may or may not be able to create the imitation of equivalent tissues requires a broad understanding of the cellmatrix interaction. Therefore, the role of the combination of materials is considered as the essential means to overcome all the barriers that include bioactivity, biodegradability, microbial contamination and maximum mechanical flexibility which contribute as a key to tissue engineering.

#### **3.1 Impact of naturally sourced biomaterials on cell-cell cross-talk**

Mimicking of biological tissue environment using collagen is an attractive theme, which is due to the inherent features such as fibrous structure, biocompatibility and low antigenicity. However, the improvement of mechanical stability and biodegradability requires additional treatment that includes cross-linking or chemical modification in the presence of second party molecules. The approaches of modification of different natural biopolymer and its tuning into desired artificial tissue architecture that support the adhesion, proliferation, and migration of cells in biological niche are discussed here.

#### *3.1.1 Collagen as base material and its derivatives*

Collagen is a key component in extracellular matrixes and composed of RGD (arginine−glycine−aspartic acid) domains that plays a potential role in cell adhesion, growth and motility through its interaction with cells. But, the drawback due to poor mechanical stability and biodegradation is overcome by the modification with various natural polymers or synthetic polymers. In one approach, collagen molecules were chemically conjugated with oxidized guar gum to immobilize platelet-derived growth factor [13]. The guar gum which is a water soluble and ionic polysaccharide was oxidized to poly(dialdehyade) guar gum in presence of sodium periodate. The resultant oxidized guar gum not only promoted the crosslinking of collagen molecules but also helped to immobilize the platelet-derived growth factor, enabling the formation of biologically active hybrid 3D scaffolds with excellent swelling, thermal and biodegradable properties. FTIR, SEM analysis was performed to confirm the synthesis of the hybrid structure. SEM morphology revealed the interconnected 3D porous honeycomb structure with an average pour size of 15 ± 7 μm. The hybrid scaffold was shown to promote the release of growth factors with the increase of NIH 3 T3 cell density and proliferation and was seen as a promising candidate for tissue engineering applications.

Recently, Diogo et al. developed a method of fabrication of '*in situ'* mineralized collagen based 3D printed hydrogel. As an alternative to various traditional approaches, Co-precipitation method is used to mineralize the collagen fiber in presence of calcium chloride (CaCl2) and ammonium hydrogen phosphate [(NH4)2HPO4]. To prepare the cell laden 3D printed hydrogel, variuos raio's of mineralized collagen and alginate (a biocompatible and degradable natural polymer) mixture was treated with incubated L929, mouse fibroblast cell line and printed using a bioprinter V1 (REGEMAT 3D, Granada, Spain), resulted the cell laden 3D printed bi-oink. The cell-laden scaffold was shown to support the adhesion, growth and survival of mouse fibroblast cell line [14]. The similar kind of the cell-laden-collagen core and alginate-polyethylene oxide shell based 3D porous structure was developed using microfluidic channel and at low temperature working condition for cryopreservation (**Figure 2A**) which is subjected to maintain the shortage of cells, tissue and organs. The *in vitro* assays of two days cryo-preserved

*Current Scenario of Regenerative Medicine: Role of Cell, Scaffold and Growth Factor DOI: http://dx.doi.org/10.5772/intechopen.94906*

#### **Figure 2.**

*(A) Application of cryopreservation/cell-printing process to tissue engineering processes. [Ref: [15], reproduced with permission from publishing authority]. (B): (a) Optical and SEM images of final scaffold with core/shell mesh structure. (b) Optical images of final scaffold before/after EDTA treatment and fluorescence images of live cells and collagen fiber in core strut after EDTA treatment. [Ref: [15], reproduced with permission from publishing authority]. (C): Osteogenic differentiation of human adipose-derived stromal cell (hASC)-laden* 

*structures with (experimental) or without (control) chicken bone marrow cell-conditioned medium. (i) Optical microscopy images of alkaline phosphatase (ALP) staining (at 7 days) and relative ALP activity (at 3, 7, and 14 days) of hASCs in the control and experimental structures. The experimental structure reached maximum ALP activities at 7 days (significantly higher than those of control) and then decreased at 14 days, whereas the ALP activities of the control group continued to increase at 14 days (n = 6, \**p *< 0.05). Decreased gray values of the magnified ALP staining, indicating that the proliferating cells in the pore were effectively mineralized. The image was captured in the pore of the experimental structure. (ii) Optical microscopy images of Alizarin Red S staining (at 14 days) and relative calcium deposition of the scaffolds (at 7 and 14 days). Calcium deposition levels of the experimental scaffolds were significantly higher than those of the control scaffolds (n = 6, \**p *< 0.05). Increased red staining was observed in the pore region in the experimental scaffold. (iii) Expression levels of runt-related transcription factor 2 (*Runx2*), collagen type I alpha chain 1 (*Col I*),*  Alp*, bone morphogenic protein 2 (*Bmp-2*), osteocalcin (*Ocn*), and osteopontin (*Opn*) at 14 days of culture. Significantly increased expression levels of* Runx2*,* Bmp-2*,* Ocn*, and* Opn *were detected in the experimental structure (n = 6, \**p *< 0.05). All values are expressed as mean ± SD. [Ref. [20], reproduced with the permission from publishing authority]. [Ref: [16], reproduced with permission from publishing authority].*

osteoblast cells or human adipose stem cells demonstrated good cell viability and steady growth similar to conventional 3D scaffold based cell treatment as shown in **Figure 2B**, which would bring the potential application in tissue engineering [15]. In a report, the mixture of neonatal chicken bone marrow cells (cBMCs) derived bioactive component and collagen was used to prepare the cell supporting bioink, which was further printed with human adipose tissue-derived stromal cell (hASC) lines to formulate the hASC-laden 3D scaffold [16]. The *'in vitro*' study using 3D architecture was shown to promote the growth, proliferation and osteogenesis of hASC cells. The system was further implanted in a rat mastoid obliteration model to monitor the potential effect of cBMCs derived bioactive component on the osteogenic differentiation in new bone regeneration. After 12 weeks of post transplantation, experimental groups showed excellent bone formation as compared as shown in **Figure 2C**. The modification of collagen fiber with synthetic polymer such as polycaprolactone or polylactic acid also resulted the mechanically and biologically active 3D porous cell supporting materials and which was confirmed by SEM, FTIR and proton NMR studies. In skin tissue engineering, the 3D scaffold was shown to increase the adipose tissue derived mesenchymal stem cell (AT-MSC) adhesion and growth with the formation of a tissue environment compared to only PCL or PLA-based scaffolds [17].

#### *3.1.2 Hyaluronic acid as base material and its derivatives*

As like collagen, hyaluronic acid (HA) is also a part of extracellular matrix and shown to have a potential role in modulating inflammation, cell attachment, and migration as well as tissue morphogenesis, owing to the biodegradable, biocompatible, non-immunogenicity and anti-inflamatory properties. An attempt by Gao et al. was initiated to develop the self-crosslinked hyaluronic acid-grafted collagen-I hydrogel using EDAC/NHS reaction method. Further, chondrocytes was encapsulated into the hydrogel to verify the cell-matrix interaction and which had a significant effect on the secretion of cartilage-specific matrices to promote the migration, proliferation and gene expression of chondrocytes cells. The *in-vivo* cytocompatibility and biodegradation studies on Sprague–Dawley (SD) rats (~200 g) showed the gradual decrease of the assembled nanofibre bundle and weight of the hydrogel after prolonged subcutaneous implantation [18]. Therefore, the manufactured self-crosslinking hydrogel could find significant application in tissue engineering applications.

In an approach, the injectable HA-SH/peptide hybrid hydrogels was developed based on the covalent/noncovalent supramolecular interaction between thiolated hyaluronic acid (HA-SH) and BPAA-AFF-OH short peptide to regulate the chondrogenic expression both *in vitro* and *in vivo* environments in the cartilage tissue

#### *Current Scenario of Regenerative Medicine: Role of Cell, Scaffold and Growth Factor DOI: http://dx.doi.org/10.5772/intechopen.94906*

engineering [19]. The prepared hydrogel confirmed by proton NMR, FTIR and SEM analysis was then employed to explore the cytocompatibility of chondrocyte cells. The chondrocyte cells laden hybrid scaffolds demonstrated the significant adhesion, proliferation of chondrocyte with the expression of chondrogenic specific genes such as Col II, Sox9 and AGG. The *in vivo* study based on the subcutaneous implantation of chondrocyte cells encapsulated hydrogel in New Zealand rabbit's models also revealed the abundant aggregation and proliferation of cells with the secretion of matrix after 4 weeks of post implantation period, resulting in the inhibition of hypertrophy trend of chondrocytes and tunable hyaline cartilage formation.

In order to improve the mechanical properties, methacrylated HA was modified with elastin-like polypeptide (ELP, consists of 70 repeats of the pentapeptide VPGVG) through free radical photopolymerization technique. The hydrogel made from the combination of MeHA/ELP revealed the tunable physicochemical and mechanical properties which is comparable to native tissue structure. Further, incorporation of zinc nanoparticles into the hydrogel had resulted an excellent antimicrobial platform for cell adhesion, growth, and proliferation phenomena (See **Figure 3A**). The '*in vivo'* cytocompatility experiment via subcutaneous implantation of MeHA/ELP hydrogel demonstrated that the weight of the hydrogel was significantly decreased with the generation and growth of autologous tissue without any inflammatory action [20]. This was due to the biodegradation of the transplanted hydrogel that led to the new space for the spreading and infiltration of the proliferated cells. The immunofluroescent staining study also exhibited the minor invasion and infiltration of lymphocyte and macrophages cells, assigning the potential application of the engineered hydrogel in various artificial tissue regeneration processes.

Like various tissue regeneration processes, utilizing of HAbased cell supporting materials explicated the significant output in artificial salivary gland repairing and remodeling. In an experiment, Lee et al. demonstrated the synthesis of hyaluronic acid−catechol (HACA) conjugates based platform named as NiCHE (nature-inspired catechol conjugated hyaluronic acid environment) to mimic the mesenchyme of embryonic submandibular glands (eSMGs) as shown schematically in **Figure 3B(i)** [21]. The NICHE was developed by the coating of HACA conjugates on the various polymeric scaffolds such as polycarbonate membrane, stiff agarose hydrogel, and polycaprolactone that led to cell adhesion and growth, vascular endothelial and proliferation of progeniotor of eSMGs cells isolated from ICR mice fetus [See **Figure 3B(ii&iii)**].

#### *3.1.3 Gelatin as base material and its derivatives*

Owing to the excellent biocompatibility, biodegradability and water solubility, gelatin has emerged tremendous interest for the formation of 3D hydrogel in tissue engineering application. Song et al. developed an injectable 3D printed gelatin hydrogel composed of continuous phase gelatin and gelatin microgels. The twostep cross-linked injectable gelatin was shown to exhibit the biocompatible lattice, cup-shaped, tube-shaped and rheological modified structure analogous to human anatomical features. The biocompatibility of the microgel led to the spread and expression of metabolic activities in mouse fibroblast cells, which can be attributed to the good cell-matrix interaction such as *in vivo* remodeling stages [22].

Due to the limited clinical success in repairing defective cartilage, unlike conventional surgery, the biopolymer-based tissue engineering approach such as the production of gelatin-linked electrospun, gelatin-polycaprolactone (gelatin-PCL) nanofiber-filled decellularized extracellular matrix has been investigated to monitor biological functions. The decellularized composite has shown to exhibit the

#### **Figure 3.**

*(A)(i) In vitro cytocompatibility of MeHA/ELP and MeHA/ELP-ZnO hydrogels. Representative live/dead images from hMSCs seeded on (a) MeHA/ELP and (b) MeHA/ELP-ZnO hydrogels after 5 days of seeding. Representative phalloidin (green)/DAPI (blue) stained images from hMSCs seeded on (c) MeHA/ELP and (d) MeHA/ELP-ZnO hydrogels at day 5 post culture. Quantification of (e) viability and (f) metabolic activity of hMSCs seeded on hydrogels after 1, 3, and 5 days of culture. Hydrogels were formed by using 2% MeHA and 10% ELP with 0 and 0.2% (w/v) ZnO nanoparticles at 120 s UV exposure time (\* p < 0.05, \*\* p < 0.01, \*\*\* p < 0.001). [Ref: [20], reproduced with permission from publishing authority]. (ii) In vitro antimicrobial properties of MeHA/ELP-ZnO hydrogels with different ZnO concentrations. Representative SEM images of methicillin-resistant* Staphylococcus aureus *(MRSA) colonization on hydrogels containing (a, b) 0% ZnO, (c, d) 0.1% ZnO, and (e, f) 0.2% ZnO. Clusters of bacteria are shown in dashed circles. [Ref: [20], reproduced with permission from publishing authority]. (B): (i/a) Schematic Diagram of Salivary Gland Damage and Tissue Engineering-Based Therapeutic Approach; (i/b) Schematic Diagram of Mimicking Mesenchymal HA of Developing eSMG on Material Surfaces by Using Adhesive HACA. [Ref: [21], reproduced with permission from publishing authority]. (ii) NiCHE coating platform enhances growth of eSMGs on various substrates.* 

#### *Current Scenario of Regenerative Medicine: Role of Cell, Scaffold and Growth Factor DOI: http://dx.doi.org/10.5772/intechopen.94906*

*(b) Bright field images and budding levels of eSMGs cultured either on bare or HACA-coated materials after 48 h-culture. Average bud count of eSMGs freshly isolated from TP15 mouse fetus is considered as 100%. Scale bar = 200 μm. (iii) Apoptotic activity, VE structure, and mitotic activity of progenitor cells of eSMGs cultured either on bare or HACA-coated materials after 24 and 48 h-culture, respectively. White scale bar = 200 μm, red scale bar = 500 μm, yellow scale bar = 100 μm. [Ref: [21], reproduced with permission from publishing authority]. (C) (i) Gross view and histological evaluation of repaired region at 6 and 12 weeks postsurgery. Macroscopic images of cartilage defects regions at 6 weeks (A*−*C) and 12 weeks (D*−*F) postsurgery. (ii) Biomechanical and biochemical analyzes of the engineered cartilage tissue in vivo. Young's modulus (A), wet weight (B), thickness (C), DNA qualification (D), GAG qualification (E), and collagen qualification (F) of the in-vivo-engineered cartilage tissue. Values are expressed as mean ± SD, n = 3, \*p < 0.05. [Ref: [23], adapted with permission from publishing authority]. (D) (i): (A1-A3) Spatial distribution of the purple formazan crystal inside the 5 L implant at day1(A1), 3(A2) and 5(A3). The scale bars represent 5 mm. Distribution profile of purple formazancrystal indicates an in-growth of cells inside the implant. (B1–5) Presence and distribution of MTT crystals on each bead layer. (ii) Study of the osteogenic properties of the implant in vitro. The study was carried out using human mesenchymal stem cells (hMSC). (A) Time dependent variation of alkaline phosphatase expression. ALP activity was measured in the supernatant of the culture. (B) Study of the expression of osteogenic marker through RT-PCR. The study was carried out after culturing the hMSC on different substrates in presence of osteogenic media for 14 days. (C) Fluorescent micrographs of the replated partially differentiated hMSC stained with FITCPhalloidin(green) and DAPI (blue). The cells were initially cultured on implant and control substrates. They were then trypsinized and replated on tissue culture plate. Imaging was done after 3 days of replating. [Ref: [26], reproduced with permission from publishing authority].*

excellent mechanical property and promoted the cartilage regeneration with the secretion of collagen and glycosaminoglycan as shown in **Figure 3C** [23].

The development of gelatin methacrylate (GelMA) and poly (ethylene glycol) diacrylate (PEGDA) printed three layered scaffold, modified by lysine functionalized rosette nanotubes (RNTK) significantly improved the adhesion, growth and differentiation of adipose-derived mesenchymal stem cells (ADSCs). The RNTK not only acted as a potential biomimetic layer, its presence dramatically increased the secretion of collagen II, glycosaminoglycan, and total collagen as compared to native GelMA-PEGDA scaffolds, and have an potential impact on cartilage regeneration [24].

To improve wound repair caused by burns or accidental injuries, a versatile approach has been shown to fabricate the skin tissue analogue of a mechanically stable acellular elastomeric scaffold in the presence of biodegradable polyurethane and gelatin composite. The Gel-20%PU showed the best cell infiltration and biodegradation in a mouse *in vivo* experiment. Also, it reveals negligible immunogenicity and could be accepted as a substitute for new generation tissues [25].

#### *3.1.4 Sodium alginate as base material and its derivatives*

The implementation of the osteogenic microenvironment loaded with therapeutic agents has emerged as the key pathway for bone tissue engineering in recent decades. Like various biopolymers, the utility of alginate, which is a polyionic-polysachharide comprising units of mannuronic acid and guluronic acid, has strengthened the field of next-generation polymer remodeling. The fabrication of calcium alginate bead based 3D implant made by the stacking of hexagonal closed pack (HCP) layers (**Figure 3D(i)**) in presence of glutaraldehyde crosslinker facilitated the spatiotemporal drug release in the artificial matrixes through the changes of the spatial coordinates of the drugs loaded layers. The supporting scaffold promoted the growth, progression and cytosketal reorganization of the osteoblast cells and triggered the expression of the alkaline phosphatase, runx2 and collagen type1 in human mesenchymal stem cells, attributed to the osteoconductive and osteogenic nature of the implant [26]. The *in vivo* assessment of the VEGF loaded implant was conducted in mice model and it revealed the regeneration of tissue with prominent existence of neovascularization as shown in **Figure 3D(iii)**, due to cohesive interaction between supportive implant and native tissue environment. A recent report [27] demonstrated the formation of mechanically stable alginate-gelatin

(ALG-GEL) hydrogels, resembling the comprehensive nonlinear and complex mechanical features of brain soft tissues. The rheology analysis also indicated that the stiffness of the hydrogel is solely dependent on the blending concentration and incubations times of the composites, assigning for the potential application in the fabrication of brain tissue supporting matrixes. Another report, where, (2, 2, 6, 6-Tetramethylpiperidin-1-yl)oxyl or (2,2,6,6-tetramethylpiperidin-1-yl) oxidanyl [TEMPO] oxidized cellulose nanofibre incorporated alginate scaffolds was investigated to enhance the biodegradability, sustainability and mechanically strengthened reactive surface area. The rheology study resulted in the recovery of 60% viscosity than that of native alginate scaffolds. The simulated body fluid (SBF) mediated mineralization evolved the nucleation of the hydroxyappetite into the hydrogel [28]. The combinatory efforts including direct writing 3D printing and freeze-drying techniques were carried out to develop the dual porous mechanically and dimensionally stable cellulose based 3D scaffolds. The dehydrothermal treatment displayed the increased surface hardness, indentation modulus and compression strength, should opened the new glimpse toward the decoration of bio-mimetic bone tissue engineering [29]. Furthermore, several studies have been conducted using cellulose or different derivatives to present an active tissue engineering tool and revealed the positive result in terms of cytocompatibility, biodegradability and flexibility of the scaffolds [30]. Similarly, Suneetha et al. [31] described the synthesis of mussel inspired polydopamine (PDA) filled sodium alginate (SA) − polyacrylamide (PDA − SA − PAM)-based hydrogel for skin tissue regeneration. The *in situ* synthesis process was conducted via two consecutive reaction steps. Initially, the dopamine molecules in the dopamine and sodium alginates blend was polymerized through alkali-induced polymerization and secondly, free radical polymerization technique was used to polymerize the acrylamide part in the processing of mechanically and biologically supporting adhesive hydrogel. The cytocompatibility assessment of the human skin fibroblasts (SFs) and keratinocytes (KTs) seeded PDA − SA − PAM-based hydrogel exhibited the higher cell adhesion, proliferation and spreading into the 3D microenvironment as compared PDA-free or 2D polystyrene plate and which is confirmed by fluorescence based live-dead assays or SEM morphology analysis as shown in **Figure 4A(i&ii)**. In addition, the effect of PDA molecules on the platelet adhesion was evaluated by the processing of porcine whole blood and it showed the higher adhesion of the platelet on the hydrogel as shown in SEM images, attributing to the potential effect of PDA in the regulation of fibrous network and adhesion of bioactive molecules.

#### *3.1.5 Chitosan as base material and its derivatives*

Chitosan, a polysaccharide with various functional groups has increased tremendous interest in biomedical applications such as tissue engineering. But, major problems due to poor solubility and biodegradability limit its monopoly use in the processing of cell supporting materials. This is avoided by the stacking or modifying with various synthetic or natural biomaterials. In a report, Li et al. developed the oxidized alginate hydrogel crosslinked with N, O-carboxymethyl chitosan with moderate swelling, degradation and porosity. The chitosan modified alginate scaffold revealed improved biocompatibility, as the number of free aldehyde groups in the oxidized alginate is reduced after crosslinking [32].

In another report [33], the methacrylated chitosan molecules were conjugated with lysozyme (an endo-carbohydrase) via riboflavin initiated photo-cross-linking to a constructed biodegradable and biocompatible hydrogel. The *in vitro* biodegradation study of the hydrogel revealed the increase of the pore size and larger fraction of outliers in cryo-SEM micrographs. Further, the mouse bone marrow

*Current Scenario of Regenerative Medicine: Role of Cell, Scaffold and Growth Factor DOI: http://dx.doi.org/10.5772/intechopen.94906*

#### **Figure 4.**

*(A) (i) Schematic Representation of the Formation of Adhesive Hydrogels for Tissue Engineering Applications. (ii) SEM images of cell attachment of SF and KT cells on hydrogels in different culture conditions (3 and 7 days). (iii) MTT assay cell proliferation of hydrogels for (a) SF and (b) KT. (iv) live/dead assay fluorescence images of SF and KT cells on hydrogels under 10× magnification (scale bar 200 μm) (\*p < 0.05). [Ref: [31], reproduced with permission from publishing authority]. (B) Schematic presentation of Chitosan-Lysozyme Conjugates for Enzyme-Triggered Hydrogel Degradation in bone Tissue Engineering Applications. [Ref: [33], reproduced with permission from publishing authority]. (C) (i) SEM images of the cell seeded scaffold of pure CHT and its nanohybrid. Red arrow indicates the position of the cells. Scale bar = 60 μm; (ii) (a) X-ray photographs of rat femur bone defects of control (devoid of any material), CHT (filled with pure CHT scaffold), and CHT-L (filled with nanohybrid scaffold, CHT-L) after one day, four weeks and eight weeks of implantation. Blue and cyan arrows indicate osteocyte and osteoblast cell, respectively; (iii/a) H & E stained histopathological section of the rat femur bone defects after eight weeks of implantation, inset figure of each image indicates that the section were taken from that part (indicated by dotted red lines). Scale bars represent 100 μm; (iii/b) Histopathological section of connective tissue attached with the bone, stained using H & E. Scale bar is 100 μm. [Ref: [35], reproduced with permission from publishing authority].*

stromal cell line (BMSC) loaded hydrogel exhibited the adhesion, proliferation and spreading of BMSCs with the expression of the osteogenic-specific markers throughout various layers of hydrogel as compared to chitosan based hydrogel. This may be attributed to the lysozyme mediated breaking of chitosan chains, thereby helping the penetration of cells through the disintegrated hydrogel networks. Also, as shown in **Figure 4B,** the *in vivo* bone regeneration experiment demonstrated the significant recovery of the defected bone with new bone tissues after six weeks of hydrogel post-implantation in a nude mice model with the recruitment of cells to the damaged zone. Similarly, the mechanically stable and tunable porous graphene oxide incorporated alginate-chitosan-collagen (GO/SA-CS-Col) based composite scaffolds was fabricated via Ca2+ mediated crosslinking and freezedrying techniques. The composition and surface morphology were confirmed by FTIR, Raman spectra, SEM and XRD analysis. *In vitro* study of the osteoblast cell encapsulated scaffold revealed the adhesion, proliferation of the cells and osteogenic differentiation [34]. In an approach, the sulfonated graphene oxide functionalized chitosan based hybrid scaffold was prepared and FTIR, SEM, TEM and XRD analysis confirmed the synthesis of the interconnected porous chitosan/ GO nanohybrid scaffolds [35]. The *in vitro* drug release assay at phosphate buffer solution at 37°C exhibited the sustained release of the drug molecules, may be due

to the noncovalent interaction between drug and composites that triggering the slow diffusion of the drug molecules. The hybrid scaffold also revealed the high cell growth and spreading into the deep part of porous scaffold as compared to GO-free chitosan scaffold which was observed by fluorescence and SEM analysis as shown in **Figure 4C(i)**. Further, the *in vivo* experiment using cell laden scaffold in rat model was investigated and the 75 days of post-implantation result demonstrated the faster healing of the defected area with significant proliferation of the osteoblast cells as compared to pure chitosan scaffold (see in **Figure 4C(ii)**). Intertestingly, Zao et al., developed a glucono δ-lactone (GDL) incorporated and carboxymethyl chitosan (CMCh) stabilized calcium phosphate (ACP) (designated as CMCh-ACP hydrogel) bioactive hydrogel using freeze-drying process for mesenchymal stem cells [MSCs] based bone regeneration [36]. FTIR analysis exhibited the characteristics peaks at 1064and 547 cm−1 due to the presence of phosphate (PO4 3−) group of spherical particles with average size of 80 nm. Next, the cytocompatability of the MCSs laden (iMAD cell line) CMCh-ACP hydrogel revealed the time dependent increase of cell density with negligible apoptotic cell morphology as in Hoechst 33258 stained, indicating the biocompatibility of the hydrogel for long-term cell friendly growth microenvironment. Like, *in vitro* assays, the *in vivo* bone regeneration experiment in presence of BMP9-(potent bone-forming factor) induced iMAD cells/CMCh-ACP hydrogel demonstrated the efficient new bone formation with the extensive vascularisation on the surface of the masses, attributing to the upregulation osteogenic-specific biomarkers and regulars, thereby enabling the BMP9 induced osteogenesis. In our several works, we demonstrated the modification of chitosan with different biomaterials such as montmorillonite clay (OMMT), hydroxyapatite, poly (ethylene glycol), polymethylmethacrylate-co-2-hydroxyethyl-methacrylate and polyvinyl alcohol. The formulated porous scaffolds were made with improved mechanical, antibacterial and biocompatible for application in bone tissue engineering [37–40].
