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Tissue Engineering & Additive Manufacturing (TEAM) Lab, Centre for Nanotechnology & Advanced Biomaterials, ABCDE Innovation Centre, School of Chemical & Biotechnology, SASTRA Deemed University, Thanjavur, Tamil Nadu 613 401, India
Tissue Engineering & Additive Manufacturing (TEAM) Lab, Centre for Nanotechnology & Advanced Biomaterials, ABCDE Innovation Centre, School of Chemical & Biotechnology, SASTRA Deemed University, Thanjavur, Tamil Nadu 613 401, India
Tissue Engineering & Additive Manufacturing (TEAM) Lab, Centre for Nanotechnology & Advanced Biomaterials, ABCDE Innovation Centre, School of Chemical & Biotechnology, SASTRA Deemed University, Thanjavur, Tamil Nadu 613 401, India
Tissue Engineering & Additive Manufacturing (TEAM) Lab, Centre for Nanotechnology & Advanced Biomaterials, ABCDE Innovation Centre, School of Chemical & Biotechnology, SASTRA Deemed University, Thanjavur, Tamil Nadu 613 401, India
Polysaccharide based hydrogels have been predominantly utilized as ink materials for 3D bioprinting due to biocompatibility and cell responsive features. However, most hydrogels require extensive crosslinking due to poor mechanical properties leading to limited printability. To improve printability without using cytotoxic crosslinkers, thermoresponsive bioinks could be developed. Agarose is a thermoresponsive polysaccharide with upper critical solution temperature (UCST) for sol-gel transition at 35–37 °C. Therefore, we hypothesized that a triad of carboxymethyl cellulose(C)–agarose(A)–gelatin(G) could be a suitable thermoresponsive ink for printing since they undergo instantaneous gelation without any addition of crosslinkers after bioprinting. The blend of agarose-carboxymethyl cellulose was mixed with 1% w/v, 3% w/v and 5% w/v gelatin to optimize the triad ratio for hydrogel formation. It was observed that a blend (C2-A0.5-G1 and C2-A1-G1) containing 2% w/v carboxymethyl cellulose, 0.5% or 1% w/v agarose and 1% w/v gelatin formed better hydrogels with higher stability for up to 21 days in DPBS at 37 °C. Further, C2-A0.5-G1 and C2-A1-G1hydrogels showed higher storage modulus 762 ± 182 Pa & 2452 ± 430 Pa, higher porosity of 96.98 ± 2% & 98.2 ± 0.8% and swellability of 1518 ± 68% & 1587 ± 25% respectively. To evaluate the in vitro potential of these bioink formulations, indirect and direct cytotoxicity were determined using NCTC clone 929 (mouse fibroblast cells) and HADF (primary human adult dermal fibroblast) cells as per the ISO 10993-5 standards. Importantly, the printability of these bioinks was confirmed using extrusion bioprinting by successfully printing different complex 3D patterns.
Owing to precise cell positioning and ability to control 3D microarchitecture, additive manufacturing of biological constructs have gained momentum in the field of regenerative medicine [
]. Among additive manufacturing techniques, 3D bioprinting offers high reproducibility and repeatability in the fabrication of complex biological tissues with better cytocompatibility [
]. In 3D bioprinting, cellular layers are stacked layer-by-layer in three dimensions by continuous printing of cell-encapsulated hydrogel bioinks. Extrusion bioprinting is a widely-utilized technique in the fabrication of 3D cellular hydrogel constructs due to its accessibility, high reproducibility and ability to print wide varieties of bioinks [
]. Bioinks used for extrusion bioprinting could be either pre-crosslinked hydrogels or non-crosslinked hydrogels embedded with cells. Although this method provides high-resolution printing with approximately 95–98% cell viability, the requirement of post-crosslinking steps complicates the fabrication process and reduces the stability of hydrogel constructs & viability of cells. Further, extrusion bioprinting suffers with limitations such as nozzle clogging, bioink experience higher mechanical stress at the nozzle tip, and poor hydrogel stability. To overcome these bottlenecks, biomaterial/bioink composition could be modified or formulated to produce novel composite bioinks with better printability, cytocompatible crosslinking choices, stability and cell viability [
]. In recent years, more progress has been made in developing natural polymer bioinks like gelatin, sodium alginate, hyaluronic acid, cellulose derivatives, etc., to make improved bioinks for 3D bioprinting applications [
]. However, developing a new bioink composition with multifunctional properties including stimuli responsiveness, biological cues and higher mechanical stability remains a major challenge.
Physically crosslinked bioinks are widely used for tissue engineering applications due to ease in crosslinking and ability to provide adequate shape fidelity to the printed structures. However, such physically crosslinked bioinks require an additional crosslinking such as chemical crosslinking, photocrosslinking and stimuli-responsive crosslinking to promote long-term mechanical stability for tissue maturation. Chemical crosslinking involves the chemical crosslinkers that aid in the formation to covalent crosslinks between polymer strands that form interconnected stable hydrogel networks. Some of the classical crosslinking chemistries include Michael addition, click chemistry, photo-polymerization, condensation reaction, etc., to form a stable hydrogel by covalent bond formation [
R. Augustine, H. Alhussain, A.A. Zahid, S.R. Ur Rehman, R. Ahmed, A. Hasan, Crosslinking strategies to develop hydrogels for biomedical applications, (2021) 21–57. doi:10.1007/978-981-15-7138-1_2.
]. The presence of unreacted chemical reagents like crosslinkers (glutaraldehyde), photo initiators (lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), monomers, and oxidizing/reducing agent may induce cellular toxicity [
]. As an alternative to cytotoxic crosslinking methods, physical crosslinking strategies involving hydrogen bonding, ionic interaction and hydrophilic/hydrophobic interactions are utilized to produce biological constructs [
]. Thermoresponsive polymeric hydrogels are now used because they can rapidly transform from free-flowing viscous liquid to stable crosslinked hydrogel when a critical temperature is reached. Thermoresponsive polymers with sol-gel transition behavior due to temperature are categorized as UCST (upper critical solution temperature) and LCST (lower critical solution temperature) [
]. For bioprinting applications, natural or synthetic polymer with the transition ability from liquid to stable gel at 37 °C (at physiological conditions) is preferable to ensure higher cell viability with better stability. Hence, LCST polymers like collagen, Matrigel®, Pluronics, poly(N-isopropylacrylamide) (pNiPAM), etc., and UCST polymers like agarose, polyacrylamide derivatives, gelatin, etc., are explored for bioprinting applications [
]. Phase transition temperature for hydrogel formation is controlled by tuning the functional groups of UCST / LCST polymers to augment hydrophilic and hydrophobic interactions, which will further improve the strength and dynamics of the thermoresponsive bioinks during the extrusion bioprinting process [
Among UCST, agarose is a naturally derived polymer composed of repeating D-galactose and D-3,6-anhydrous-galactopyranose unit connected with α-1,3 and β-1,4 glycosidic bonds. Recently, agarose-based bioinks have been widely explored in the fabrication of stable biological constructs. It is a well-known biocompatible polymer extensively used for hydrogel preparation, 3D biofabrication, and slurry preparation for scaffold-free bioprinting applications [
E. Mirdamadi, N. Muselimyan, P. Koti, H. Asfour, N. Sarvazyan, Agarose slurry as a support medium for bioprinting and culturing freestanding cell-laden hydrogel constructs, https://Home.Liebertpub.Com/3dp. 6 (2019) 158-164. doi:10.1089/3DP.2018.0175.
]. In 3D bioprinting, the desired properties of bioinks such as high viscosity, instantaneous gelation, cytocompatibility and stability were achieved by the physical blending of agarose with other natural polysaccharides, proteins or synthetic polymers. However, a major drawback of agarose is the lack of cell recognition motifs such as RGD, impairing its widespread applications in tissue engineering. To improve the biological functionalities, agarose bioink may be mixed with other polymers containing cell binding motif [
]. Additionally, the agarose concentration directly affects hydrogel viscoelastic properties and stability. To tune the mechanical stability and flow behavior of the bioinks for extrusion printing, several additives are researched to develop composite agarose bioink. Nadernezhad et al., developed a nanosilicate-based nanocomposite agarose hydrogel containing Laponite® nanoparticles. Here, the addition of Laponite® increased the bioink viscosity, hydrogel compressive strength and stability, indicating the strong physical interactions between the Laponite® and agarose. Likewise, agarose composite bioinks are now utilized for biofabrication applications [
]. As described by Dravid et al., the sol-gel transition of agarose blend with thermoreversible gelatin formed a stable composite hydrogel due to the presence of hydrogen bonding and interactions between double-strand helices of agarose chains at a temperature below 25 °C. Agarose/gelatin (AG-Gel) bioink possesses shear thinning behavior, which helps in cell protection from shear damage that occurs during extrusion printing process [
Biofabrication of volumetric tissues with biomimetic tissue functionalities is very crucial for successful applications in regenerative medicine, drug testing and in vitro models. Among the various natural polymers, agarose-based composite bioinks offer better feasibility for extrusion printing at physiological conditions. Carboxymethyl cellulose (C) is a Food and Drug Administration (FDA) approved cellulose-derivative biomaterial widely explored for tissue engineering and bioprinting applications due to its viscoelastic nature, tunable mechanical property, printability, cytocompatibility and biodegradability [
]. Carboxymethyl cellulose can be blended with other bioink materials such as agarose to improve viscoelastic properties and thermoresponsive crosslinking ability that allows stable hydrogel by semi-interpenetrated network formation. However, a blend of C-agarose lack cell binding motifs could be complemented with the addition of gelatin, an ECM mimetic protein derived from collagen, which is known to promote cell attachment, migration and proliferation [
]. Hence in this work, a triad of carboxymethyl cellulose-agarose-gelatin (C-A-G) bioink was formulated and characterized for thermoresponsive gelation, stability, printability and cytocompatibility.
2. Materials and methods
2.1 Materials
Agarose low EEO (36601) and high viscosity 400–800 CPS carboxymethyl cellulose sodium salt (59,938) was purchased from SRL, India. Gelatin type A (from porcine skin, G2500) was purchased from Sigma, USA. For in vitro studies, NCTC clone 929 mouse fibroblast cells were purchased from NCCS, Pune and primary Human Adult Dermal Fibroblast (HADF) cells purchased from Himedia, India (CL005) were used for cytotoxicity studies. Dulbecco's modified eagle medium (DMEM) high glucose (AL007A, Himedia), Fetal Bovine Serum (FBS) (2027-04, Gibco), Penicillin streptomycin (15070-06, Gibco), 1 × 0.05% Trypsin EDTA (25300-62, Gibco), MTS cell proliferation assay kit (ab197010, Abcam) and other chemicals of analytical grades were used for the experiments.
2.2 Hydrogel preparation and tube inversion test
The physical blend solution was prepared by mixing individually dissolved carboxymethyl cellulose (C), gelatin (G) and agarose (A) in double distilled water. Briefly, 2% (w/v) carboxymethyl cellulose solution was stirred overnight to ensure complete dissolution. Gelatin stock solution of different concentrations of 1% (w/v), 3% (w/v) and 5% (w/v) were prepared and stirred at 45 ± 2 °C for 2 h. C2-G blend was prepared by continuous stirring for 2 h at 45 ± 2 °C to ensure complete mixing. The prepared C-G solution with different gelatin concentrations were labelled as follows: C2-G1, C2-G3 and C2-G5. The agarose stock solution of 0.5% (w/v), 1% (w/v) and 1.5% (w/v) were prepared by adding the agarose powder to double distilled water. Further, the agarose mixture was heated above 100 °C (in microwave oven) for 1–3 min and then added to the C-G blend solution at 45 ± 2 °C. The C-A-G solution was thoroughly mixed by manual pipetting in a glass vial at 37 °C for the tube inversion test. C-A-G composite solution was incubated at room temperature to confirm the gel formation by inverting the vials and photographs were taken.
2.3 Scanning electron microscopy (SEM) analysis
The hydrogel samples were freeze-dried (Alpha 2–4 LDplus, Germany) at -60 °C for 36 h before further characterization. The cross-sectional surface of the sliced hydrogels were observed under a scanning electron microscope (SEM) to study the surface morphology. Before SEM analysis, the samples were uniformly sputter-coated with gold. SEM images were acquired at 200 X and 500 X magnifications using TESCAN Vega 3 at the 10 kV voltage and 10 mm working distance [
Freeze-dried hydrogel samples (C-A-G) and pristine polymer powders (agarose, carboxymethyl cellulose and gelatin) were analyzed by FTIR to study the chemical compositions. FTIR spectra of the samples were recorded over the scan range of 4000–500 cm−1 by averaging 10 scans using a Fourier-transform Infrared Spectrometer (Spectrum 100, Perkin Elmer, USA) [
Rheological characterization of the agarose composite hydrogels (n=5) were performed using Anton Paar rheometer (MCR 302) using a 25 mm diameter parallel plate with a 1 mm gap distance at 37 °C. Further oscillatory frequency sweep test was performed at 1–100 rad/s frequency rate to calculate storage modulus (Gʹ), loss modulus (Gʺ) and loss factor (tan δ). The gelling temperature of composite hydrogels was analysed by performing a temperature sweep rheological analysis using Elastosens Bio2 (Rheolution Inc). Briefly, 3 mL of composite solutions (n=4) were loaded into a calibrated sample holder. The analysis was performed with the temperature sweep from 40 to 20 °C with 2 °C step temperature, 15 min soak time and 30 s temporal step for 2 min.
2.6 Thermal analysis
The thermal stability profile of the C-A-G composite hydrogel and pristine polymers (agarose, carboxymethyl cellulose and gelatin) were analyzed by measuring the mass loss occurrence at increasing temperature by thermogravimetry analysis (TGA, SDT Q600 -TA Instruments, Germany). The weighed samples of about 6–9 mg were placed in the aluminium pan and heated at the rate of 20 °C/minute with heating rate from 0 to 700 °C.
2.7 Porosity measurement
The porosity of the freeze-dried hydrogel samples (n=4) were measured using the liquid displacement method [
]. Initially, weighed (Ws) samples were immersed in ethanol (displacement liquid) for 30 min to ensure complete solvent penetration into the freeze-dried hydrogel. Ethanol-filled tubes with or without scaffold were weighed after complete removal of the immersed scaffolds and noted as Wa (ethanol-filled tube weight), Wi (ethanol-filled tube after scaffold immersion), and Wr (ethanol-filled tube weight after removing scaffold). From these weights, the percentage porosity was calculated from the formula
2.8 Swellability study
The swellability of the hydrogel by water uptake was studied by allowing the freeze-dried samples (n=4) to swell in Dulbecco's phosphate buffered saline (DPBS, pH 7.4) at 37 °C. Freeze-dried hydrogel samples were weighed initially and incubated in the DPBS solution for a period of 1, 5, 30 min, 1 and 6 h. At the end of incubation, swollen samples were weighed after gently removing the excess DPBS. The formula mentioned below was used for analyzing swelling percentage,
where Ws is swollen hydrogel weight and Wi is the initial weight of freeze-dried sample.
2.9 In vitro degradation study
In vitro degradation profiles of the freeze-dried C-A-G samples (n=4) was performed by gravimetric method at constant physiological pH and temperature of 37 °C. To measure the weight loss percentage, the initial weighed samples were incubated in DPBS for 1, 3, 5, 7, 14, and 21 days. The DPBS was replaced every alternate day during the study period. At the end of each time point, the samples were freeze-dried and weighed to calculate the average weight loss. The formula used was,
where Wi is the initial weight of the freeze-dried hydrogel sample and Wd is the final weight of the freeze-dried sample after the predefined incubation period in DPBS.
2.10 In vitro studies
2.10.1 Cell culture
NCTC clone 929 cells and HADF cells were used to evaluate the cytotoxicity of the C-A-G composite hydrogels. Briefly, the cells were separately cultured in a T-25 flask by supplementing DMEM cell culture media containing 10% FBS and 1% penicillin-streptomycin. The cultured flask was maintained in a CO2 incubator with a 5% CO2 supply at 37 °C for cell growth. After reaching 80–85% confluency, the cells were washed with 2 mL sterile DPBS and trypsinized using 0.05% trypsin/EDTA solution. 0.25 % trypsin/EDTA solution was used for trypsinization for human adult dermal fibroblast. Cell pellets were obtained after centrifugation at 300 x g for 5 min at 4 °C. Further, cell count was performed in haemocytometer using trypan blue exclusion method.
2.10.2 Indirect cytotoxicity analysis
The cytotoxic profile of the C-A-G composite hydrogels was evaluated using ISO 10993-5:2009 guidelines “biological evaluation of medical devices – Part 5 Tests for in vitro cytotoxicity” to perform extract cytotoxicity test [
]. Initially, C-A-G composite hydrogels were prepared under sterile conditions and incubated with DMEM supplemented with 10% FBS and 1% penicillin-streptomycin. After 24 h of incubation in CO2 incubator at 37 °C, the sterile extract was collected and used directly (E4) or in diluted form (E1-E3) to study its cytotoxic effect. The extract was diluted with freshly prepared cell culture media to obtain E1 (100% v/v), E2 (50% v/v), E3 (25% v/v) and E4 (12.5% v/v). These extracts were treated with NCTC clone 929 cultured into sub confluent cell monolayer and HADF cell monolayer for cytotoxicity evaluation. Liquified phenol of 1.5% (v/v) treated cells were considered as positive control and cell culture medium treated TCPS (tissue culture grade polystyrene) cultured cells were considered as negative control for the evaluation of toxicity. Briefly for both the cell types, the cells were seeded into 96 well plates at 10,000 cells/well and cultured for 24 h. After 24 h, the media was removed and washed with DPBS. Further, the cells were treated with E1, E2, E3 and E4 for 24 h for qualitative and quantitative assessments. After treating the cells with extracts, the cultured cells were evaluated microscopically to study the morphological changes and grade the extract effects from cytotoxicity levels 0 to 4 as per the ISO. Extracts showing only intracytoplasmic granules and the absence of other cell damages are considered to be level 0. Morphological changes and growth inhibition are considered from level 1 to level 4. Extracts inducing morphological changes in 20, 50 and 70% of the cells are categorized into levels 1, 2 and 3, respectively. Extracts causing complete disruption of cells are considered level 4.
2.10.3 Direct contact cytotoxicity analysis
As per the guidelines of ISO 10993-5:2009, direct contact cytotoxicity analyses were performed using cultured NCTC clone 929 and HADF cells Initially, agarose composite hydrogels of size 6 mm diameter and 5 mm thick were prepared as mentioned in section 2.2. Concurrently, 50,000 cells/well cell density were cultured in 24 well plates to form a subconfluent layer. After 24 h, culture medium was removed and washed with DPBS. Prepared hydrogels were placed directly in the center of the wells with the cell contact area maintaining at least 1/10th of the subconfluent monolayer. Upon direct culturing for 24 h, cell morphological changes and grading were analyzed using microscopical evaluation as mentioned in Section 2.10.2.
2.10.4 Live/dead staining and MTS assay
Live/dead staining was performed for the indirect cytotoxicity assessment. Live and dead cells were stained with 1 µL/mL of calcein-AM and 2 µL/mL ethidium bromide staining solution after washing with DPBS [
]. After incubation, the staining solution was aspirated and fresh DPBS was added. Images were obtained using a fluorescence microscope (Olympus CKX53, Japan).
For the indirect and direct contact cytotoxicity analysis, MTS assays were performed for quantitative assessments. In brief, treated wells were washed with DPBS and incubated with MTS reagent as per the established protocols [
]. 10 µL MTS reagent and 90 µL serum-free DMEM were added to each well and incubated in the dark for 4 h. After incubation, the absorbance values were measured at 490 nm using a multi-mode reader (Infinity 200 M, Tecan, USA). The cell viability percentage was calculated from the formula
2.11 Printing of C-A-G composite biomaterial ink
A1-C2-G1 bioink was printed using 3D Bioplotter (EnvisionTEC, Germany) to determine the printing potential of this bioink to make complex shapes. The prepared bioink was loaded into low-temperature cartridges and attached to a low-temperature print head maintained at 37 °C to extrude the bioink without clogging the needle. The platform temperature was maintained at 10 °C, and 5 s wait time was given before printing the next layer to promote shape fidelity by rapid thermoresponsive gelation of the extruded bioink. Bioink was extruded using a 0.4 mm needle inner diameter (NID) with 0.5 s pre-flow and post-flow at a pressure of 0.4 bars and a printing speed of 15 mm/s. Complex shapes such as clover, fan, the letter “O”, grooves, and stepwise patterns were designed in 3D modelling software (TinkerCAD) and printed using the developed bioink without cells to check its printability.
2.12 Statistical analysis
All the results were mentioned as mean ± standard deviation. For porosity evaluation and direct contact cell viability study, unpaired t-test was performed to study the significance at p <0.05. Two-way ANOVA was used to evaluate degradation at p < 0.05. One way ANOVA was used to evaluate pore size and cell viability studies at p < 0.05.
3. Results & discussion
In recent times, composite hydrogels are finding more applications in bioprinting & tissue engineering due to options in incorporation of different polymeric functionalities by reinforcing nanofillers or polymer blends through physical or chemical interactions [
]. Based on this crosslinking interaction involved, polymer composite hydrogels are categorized into interpenetrated network (IPN) and semi-interpenetrated (semi-IPN) network hydrogels. For IPN hydrogels, a homogeneous interconnected network connection is formed between two or more differently crosslinked polymer chains. IPN hydrogels have tunable mechanical properties due to the composition of polymers involved in crosslinking [
]. Specifically for semi-IPN hydrogels, intermolecular interactions are present between non-crosslinked polymeric chains with at least one crosslinked polymer chain [
]. Semi-IPN provides advantages of incorporating cell binding motif, stimuli responsiveness, antioxidant and antibacterial properties to a composite hydrogel system in addition to tuning mechanical properties [
]. Both IPN and semi-IPN types bioinks could satisfy the ideal properties such as rapid crosslinking ability, porosity, printability, stability, cytocompatibility and biodegradability [
]. The major focus of this work is to fabricate a bioink triad of carboxymethyl cellulose, agarose and gelatin composites for the fabrication of soft tissue constructs. Carboxymethyl cellulose is widely used as binders, thickening agents, gelling and emulsifiers [
]. Hence, 2% carboxymethyl cellulose is used in bioink formulation to attain the required viscosity of the C-A-G bioink that allowed smooth strand extrusion through the printing nozzle. By combining carboxymethyl cellulose and agarose with gelatin blend improved the bioink properties namely stability, cytocompatibility, printability and shape fidelity. Table 1 shows the notations for all the tested bioink formulations.
Table 1Hydrogel composites developed using various concentrations of carboxymethyl cellulose (C), agarose (A) and gelatin (G)
Agarose hydrogels are majorly known to provide thermoresponsive gelation, higher stability and biocompatibility. Incorporating thermoresponsiveness to the bioink composition improves rapid gelation kinetics and tunable structural stability without need for any chemical crosslinkers. Cambria et al., developed an elastic agarose-collagen composite hydrogel with a tunable storage modulus of 13–16 kPa. This hydrogel was used to study the cellular mechanotransduction mechanism and cell-matrix interactions [
]. Likewise, Dieng et al., fabricated a printable semi-synthetic hydrogel prepared using a physical blend of Pluronic F127/Gelatin methacrylate (GelMA). This uses a multi-step gelation mechanism of thermoresponsive physical gelation at 37 °C for making printable bioink and UV-induced photochemical crosslinking to form stable 3D constructs. The formulated bioink was initially kept at solution state under 4 °C to ensure uniform cell mixing and later formed stable gel at 37 °C due to ionic interactions. The storage modulus of the first physically crosslinked hydrogel was observed to be 3 ± 0.4 kPa which improved printability. The storage modulus was found to be further increased to 6 kPa by photocrosslinking which is required to enhance the stability. The biocompatibility of the Pluronic F127/GelMA hydrogel performed with human adult dermal fibroblast also confirmed that the printed constructs had higher cell viability and proliferation at day 14 [
]. These Thermoresponsive hydrogels could be effectively utilized for in situ bioprinting and injectable hydrogel preparation due to its instantaneous gelation kinetics [
]. Mokhtarzade et al., have used a gelatin methacrylate/agarose blend to form a multilayer gradient based injectable scaffolds for osteochondral repair. Dual crosslinking method involving agarose based physical networking at UCST promote hydrogel injectability and photocrosslinking by UV light of 365 nm wavelength with 9 mW/cm2 intensity produce a stable hydrogel scaffold. Four-layer scaffold with gradient porosity of 75.84 ± 3.79% to 95.17 ± 3.47% was fabricated by varying the agarose (5 to 10 % w/v) and gelatin methacrylate (85 to 100 % w/v) concentrations. RGD binding motif present in gelatin methacrylate facilitate cell binding and intrinsic properties of agarose hydrogel promoted osteochondral regeneration. Extract cytotoxicity evaluation using human adipose mesenchymal stem cells on day 7 for four different hydrogel layers confirmed the higher cell viability of 86 to 91%. This confirmed the potential application of GelMA/agarose hydrogel to produce in situ forming injectable hydrogel scaffolds for osteochondral repair [
]. Based on the existing literature on thermoresponsive bioinks, it may be observed that the thermoresponsive triad bioink composition of carboxymethyl cellulose-agarose-gelatin prepared in this study will find potential applications in tissue engineering.
3.1 Agarose composite hydrogel formation and morphological characterization
A semi-interpenetrated (semi-IPN) network of stable C-A-G composite hydrogels were successfully fabricated by lowering the temperature 24-25 °C due to the crosslinking of single/double strand helical structures present in the agarose chains through hydrogen bonding. Agarose polymer concentration and temperature directly influenced the crosslinking mechanism and stability. Upon cooling the high-temperature aqueous agarose solution, a change in secondary structure occurs from random coil to single and double- helices. These helices aggregate and bundle to form crosslinked fibrillar networks [
]. Tube inversion tests were carried out to confirm the stable hydrogel formation by maintaining the bioink solutions at room temperature for 2 minutes. Upon incubation, bioinks containing agarose formed stable opaque hydrogels at room temperature, whereas the C-G blend without agarose used as control failed to form a hydrogel. The photographs of the inverted glass vials in Fig. 1A(i) confirmed the gelation of C2-A0.5, C2-A1 and C2-A1.5 hydrogels with different gelatin concentrations (1, 3 and 5%). The tube inversion method results demonstrated that the thermoresponsive behaviour of agarose was unaffected by the presence of gelatin and carboxymethyl cellulose in the composite bioink.
Fig. 1[A] Tube inversion method. (i) Photograph of the inverted glass vials confirming hydrogel formation in C-A-G composite (dashed yellow line) while C-G composite fails to crosslink (dashed red line). (ii) Photographs of the inverted glass vials incubated at room temperature (24-25°C) confirming the hydrogel formation of gelatin (1, 3 and 5%) blended with different combinations containing carboxymethyl cellulose and agarose. [B] 100X (scale bar-200 µm) and 500X (scale bar-100 µm) magnification SEM images of the cross-sectioned composite hydrogels. [C] Pore size calculated from SEM images of C2-A0.5, [D] C2-A1 and [E] C2-A1.5 composite hydrogels with varied concentrations of gelatin. (*p < 0.05), (ns> 0.05).
The morphology of the C-A-G hydrogels was analyzed using SEM, which showed porous architecture with interconnected network formation. For increased agarose concentration, the distribution of pores was found to increase. Low agarose concentration (0.5%) had more pores while 1.5% agarose composite had lesser pores. Similar variation in the pore distribution with decreased pore size for lower agarose concentration was observed for all the gelatin blends (1, 3 and 5%). For 1% gelatin composites, the pore size increased from 123 ± 27 µm to 300 ± 62 µm when agarose concentration was increased from 0.5 to 1.5%. Quarta et al., have identified the porosity variation in the agarose-collagen composite hydrogels with an increased porous structure for decreasing agarose concentration [
]. However, in all the agarose concentrations with varied gelatin compositions are known to affect the pore size distribution. With higher gelatin concentrations, presence of higher C-G entanglement affected agarose polymer chains mobility and crosslinking ability, which resulted in bigger pores. This was observed in blends of hydrogels containing 0.5% agarose and 2% carboxymethyl cellulose, which displayed increased pore size and pore wall thickness with increasing gelatin content from 1 to 5% (Fig. 1B(i–iii)). Based on the Image J analysis for pore size calculation, an increase in gelatin concentration of 1%, 3% and 5% in 0.5% agarose hydrogel composites resulted in an increased pore size of about 123 ± 27 µm, 132 ± 30 µm and 300 ± 62 µm respectively (Fig. 1C). A similar pattern in pore size distribution was also observed in the hydrogels prepared using 1 and 1.5% agarose. For 1% agarose composites, pore sizes were 215 ± 61 µm, 255 ± 52 µm and 380 ± 76 µm for gelatin 1, 3 and 5% concentrations respectively (Fig. 1D). For 1.5% agarose, pore sizes were in an increasing pattern with an increase in gelatin concentration as follows: 300 ± 62 µm (1% gelatin), 380 ± 76 µm (3% gelatin) and 680 ± 193 µm (5% gelatin) (Fig. 1E).
The results observed in the present study are similar to previous reports by Gong et al., in which they found that the agarose, gelatin and alginate ratio affected the pore morphology of the composite hydrogels. Increasing gelatin concentration from 5% to 15% led to larger pores with increased wall thickness in alginate/gelatin/agarose hydrogels [
Interestingly, all C-A-G composite hydrogels showed highly interconnected porous networks. However, the pore diameters of composites containing 1.5% agarose were larger compared to composites containing 0.5 and 1% agarose. This may be attributed to the fact that presence of C-G entanglements influences the agarose chain aggregations at higher concentrations, which result in increasing the pore size. Usually, these micron-sized pores are known to facilitate cell attachment, migration, nutrient diffusion and microcapillary formation (neovascularization). Manual pipetting of bioink, which contained 1.5% agarose was difficult due to increased viscosity. Hence, to minimize pipetting errors and to improve the repeatability, C2-A0.5 and C2-A1 bioink compositions were chosen for further characterization.
Intermolecular interactions between semi-IPN hydrogels of carboxymethyl cellulose-agarose-gelatin were studied using FTIR- peaks. All the peaks present in carboxymethyl cellulose, agarose, gelatin were also observed in the spectrum of composite hydrogel (Fig. 2) and tabulated in table 2 representing their corresponding functional groups. Briefly, broad absorption bands at 3700–3000 cm−1 represent the presence of O-H and N-H stretching vibration in pristine carboxymethyl cellulose, agarose, gelatin and its composites [
]. For composite hydrogel, the O-H stretching peak showed a shift from 3477 cm−1 to 3416 cm−1 compared with pristine agarose. This shift confirmed the presence of intermolecular interactions like hydrogen bonding and van der Waals interactions of carboxymethyl cellulose-gelatin in the agarose blend [
Among the properties of hydrogels, hydrophilic environment with ideal viscoelastic properties is very crucial in mimicking the native tissues. To understand the viscoelastic properties of the hydrogel, rheological studies were performed with known stress and strain values to determine the hydrogel mechanical behavior. Storage modulus (Gʹ) provide the details about hydrogels structural stability by measuring the stored energy during the deformation process. Besides, loss modulus (Gʺ) measurement determines the dissipated energy during the deformation [
]. Gʹ represents the hydrogel strength/elastic behavior, while Gʺ represents the viscous/ flow behavior. The oscillatory frequency sweep test was performed to determine the strength of hydrogels by calculating the storage modulus (Gʹ), loss modulus (Gʺ) and loss factor (tan δ) values. Generally, for a strong hydrogel system, the storage modulus should be higher than the loss modulus (Gʹ > Gʺ), which indicates their higher structural integrity properties [
]. The loss factor (tan δ) value is the ratio of the loss modulus to the storage modulus, which is commonly used to determine the hydrogel behavior (elastic or fluid). The tan δ value more than one (tan δ >1) indicate the fluid behavior of the sample, while tan δ <1, denote the steady and stable gels formation. In this study, for both C2-A0.5 and C2-A1 composites with 1% and 3% gelatin, Gʹ values were greater than Gʺ representing the stable hydrogel formation. Fig. 3A shows that at a lower gelatin concentration of 1%, an increase in agarose concentration resulted in increased storage modulus from 762 ± 182 Pa to 2452 ± 430 Pa. At 1% A, the number of crosslinking sites between the helical structures would be higher compared to 0.5% A. Thus, higher agarose concentration promotes the formation of semi-IPN hydrogels with better rigidity and high interconnectivity. Similar results of increased storage modulus with increase in agarose concentration were observed by Zhang et al., during the fabrication of silica or hectorite-reinforced agarose composite hydrogels (SA or HA). When agarose concentration was increased from 0.1% to 0.4% in silica and hectorite, hydrogels storage modulus was increased from 561 to 782 Pa and 3478 to 4700 Pa [
]. Further, for 3% gelatin, C2-A0.5-G3 and C2-A1-G3 composite hydrogels showed lower gel strength of 149.8 ± 114 Pa and 355.5 ± 52 Pa, respectively. Increase in gelatin content decreased the hydrogel mechanical strength due to the presence of higher carboxymethyl cellulose-gelatin entanglements which limited agarose crosslinking. Likewise, for 5% gelatin concentration, a crossover of Gʹ and Gʺ was observed in both the C2-A0.5-G5 and C2-A1-G5 composites indicating the occurrence of liquid behavior of hydrogels at 37°C (Fig. 3C and F). These results indicate that when gelatin concentration is increased, the mechanical stability of agarose composite hydrogels are decreased due to the inhibition of agarose crosslinking mechanism.
Fig. 3Storage modulus (G') and loss modulus (G") of composite hydrogels C2-A0.5 [A, B & C] and C2-A1 [C, E & F] with different gelatin concentrations (1, 3 and 5%).
To further confirm this result, tan δ were determined for all the composites. From Fig. 4A and B, the tan δ values of C2-A0.5-G5 and C2-A1-G5 were found to be nearly 1, stating that the composite hydrogel exhibited with dominant liquid-like behavior. This liquid-like behavior of bioink was unsuitable for bioprinting since it results in poor shape fidelity of printed structures. Recently, Dravid et al., have observed an increase in tan δ value reaching 0.42 ± 0.214 for 2% agarose hydrogel containing 1.5% gelatin. However, for lower concentration of 0.5 to 1% gelatin, tan δ value was observed to be within 0.25 to 0.45. At higher gelatin concentrations, liquid like properties were observed [
]. From these results, C2-A0.5-G1 and C2-A1-G1 composite hydrogels had a higher G' value equivalent to soft tissues or organs like kidney, brain, heart, etc., [
]. Hence C2-A0.5-G1 and C2-A1-G1 composite hydrogels were analysed further, including water absorption ability, porosity and in vitro degradation to optimise the bioink concentration.
Fig. 4Loss factor (tan δ) values of three different gelatin concentrations of 1, 3 and 5% for [A] C2-A0.5 and [B] C2-A1 composite. [C & D] shows the temperature sweep analysis results of C2-A0.5-G1 and C2-A1-G1 hydrogels. [E] TG-DTA graph of pristine agarose, carboxymethyl cellulose, gelatin and composite hydrogels.
Further, to study the thermoresponsive gelling kinetics of agarose composite blends, a temperature-dependent rheology test was performed ranging from 20 to 40 °C to determine the gelling temperature (Tg). This could help to maintain bioink suspensions in liquid state for extrusion bioprinting. Fig. 4C shows the occurrence of C-A-G composite hydrogel sol-gel transition from 36 to 26 °C. This sol-to-gel transition of C2-A0.5-G1 hydrogels leads to an increase in storage modulus from 499 ± 42 Pa at 36°C to a maximum of 2472 ± 417 Pa at 20 °C. Similarly, C2-A1-G1 hydrogels showed the sol-gel transition with increased storage modulus were observed at 36 °C with G' of 492 ± 33 Pa and 8714 ± 776 Pa at 20 °C. To better understand this phase transition, the change in loss factor value was plotted against temperature (Fig. 4D). Increase in tan δ value was observed at 34 and 36 °C for C2-A0.5-G1 and C2-A1-G1 respectively, indicating the physical change from sol-to-gel. An increase in agarose concentration increased the sol-gel transition temperature for the composite hydrogels.
Thermal analysis using TG-DTA help to understand the thermal degradation behavior of C-A-G composite hydrogel compared to the pristine polymers (Fig. 4E). Agarose and gelatin have two decomposition stages, with the initial stage of weight loss between 35 and 100 °C and the second stage between 250 and 450 °C. In comparison, carboxymethyl cellulose was observed to possess three degradation stages: the initial stage between 35 and 100 °C, the second stage between 150 and 230 °C and the third stage between 250 and 350 °C. The initial weight loss in all the samples occured near 100 °C due to the removal of moisture content. The second stage of degradation for agarose and gelatin had a weight loss percentage of 20–74% and 15–75%, respectively, due to the decomposition of their organic components [
Shapable bulk agarose–gelatine–hydroxyapatite–minocycline nanocomposite fabricated using a mineralizing system aided with electrophoresis for bone tissue regeneration.
Synthesis of cross-linked carboxymethyl cellulose and poly (2-acrylamido-2-methylpropane sulfonic acid) hydrogel for sustained drug release optimized by Box-Behnken Design.
]. For C2-A0.5-G1 and C2-A1-G1 composite hydrogels, similar weight loss profile was observed at initial stage due to the removal of bounded water molecules. Further, higher weight loss occurred in composite hydrogels at 250–400 °C, similar to pristine A, C, and G due to the decomposition of polysaccharide chains (agarose & carboxymethyl cellulose) and gelatin [
]. The thermal stability of the composites was improved by combining the properties of agarose, gelatin and carboxymethyl cellulose indicating the strong interactions between the composites.
3.3 Porosity and swelling
Liquid displacement method was used to evaluate internal pore distribution in the composite hydrogels, where ethanol as a displacement liquid penetrates through the pores. The calculated porosity of A0.5-C2-G1 and C2-A1-G1 hydrogels were 96.98 ± 2% and 98.2 ± 0.8% respectively (Fig. 5A). Highly porous hydrogels support high water absorption & nutrient diffusion, improve cell migration and promote cell-cell communication. Qi et al., had fabricated a tunable porous hydrogel with a pore diameter of 19-77 µm by varying the ratio (0:1, 2:8, 4:6 and 6:4) of 2 w/v% agarose-2 w/v% salecan. This varied pore size directly affected swelling ratio and mechanical strength of the gels. Highly porous 6:4 hydrogel had a higher swelling ratio of 71.7 and a lower gel strength of 561 Pa where the large pore promoted better water diffusion [
]. The swelling percentage determines the water-absorbing ability of hydrogels in hydrophilic physiological conditions. For both 0.5 and 1% agarose composite hydrogels, similar swelling profiles were absorbed with the maximum swelling percentage of 1647 ± 82% and 1517 ± 166%, respectively (Fig. 5B). Both hydrogels with 0.5 and 1% agarose attained equilibrium swelling after 60 min. Though a significant difference was not present, 0.5% of agarose concentration showed a higher swelling rate than 1% agarose in C2-G1 composite gels. Roberts et al., have determined the effect of swelling ratio for varied agarose concentrations of 1, 5, 10 and 15%. A significant change in the hydrogel swelling behavior was observed at higher concentrations of agarose. Based on the literature review, the agarose-based hydrogels are reported to reach their swollen state in one hour [
The degradation study showed a steady progressive degradation profile for the C-A-G blends (Fig. 5C). Fig. 6A represent the photographs of 0.5% and 1% agarose composite hydrogels at different time points. Initially, huge weight loss 49.4 ± 2% and 48.2 ± 4% were observed in both C2-A0.5-G1 and C2-A1-G1 blends respectively at day 3. This higher degradation was due to the release of physically blended gelatin and carboxymethyl cellulose from the agarose composite hydrogels. C2-A0.5-G1 hydrogel (69.6±4%) showed significantly higher degradation compared to C2-A1-G1 hydrogel (60.3 ± 2%) on day 21. This may be due to the intrinsic property of agarose to offer higher stability to the composite hydrogels [
]. Rossi et al., have obtained a similar higher degradation behavior of 53% weight loss at day 28 for sodium fluorescein-loaded agarose-carbomer hydrogels (AC1_SF) [
Fig. 6[A] Photographs and [B] Cross sectional SEM images of C2-A0.5-G1 and C2-A1-G1 hydrogel samples incubated in DPBS for various time points (200 X magnification with scale bar: 200 µm and 500 X magnification with scale bar: 100 µm (inserts)).
Fig. 6 represents the cross-sectional SEM images of hydrogel samples incubated in DPBS for various time points. It was observed that with increase in incubation time the pore size of composite hydrogels were increased due to degradation. In longer time points (day 14 and day 21) pore morphologies started to disappear along with the appearance of fibrous structures. These results further confirm the degradation behavior (Fig. 5C).
3.5 Indirect cytotoxicity study
Bioink used for biofabrication should be cytocompatible to facilitate cell-fate processes and augment tissue regeneration. To understand the cytotoxicity effect, an indirect cytotoxicity assay was performed for C2-A0.5-G1 and C2-A1-G1 semi-IPN hydrogels as per ISO 10993-5:2009. NCTC clone 929 cells (mouse fibroblasts) and human adult dermal fibroblasts, were used to evaluate the cytocompatibility of the developed.
In this experiment, microscopic evaluation of cell response with a grade 2-4 and cell viability of less than 70% is considered cytotoxic. According to the ISO, materials showing the above response should be avoided for biological applications [
]. In Fig. 7A (NCTC clone 929 cells), all the extract dilutants of 0.5% agarose composite hydrogels (C2-A0.5-G1) like 100%, 75%, 50% and 25% were found to exhibit cell viability of 103 ± 10%, 85 ± 6%, 88 ± 6% and 87 ± 3%, respectively. Similarly,100%, 75%, 50% and 25% extracts of 1% agarose composites resulted in cell viability of 97 ± 11%, 89 ± 6%, 82 ± 8% and 98±14%, respectively. These results confirm the non-cytotoxicity of the C2-A0.5-G1 & C2-A1-G1 hydrogels and are suitable for biofabrication of tissue constructs. Further morphological evaluation of the NCTC clone 929 cells cultured using 100% extract confirmed native spindle-shaped morphology with a higher extension as in the case of TCPS control group. While in the positive control (phenol-treated group), morphological changes with the uneven cell surface and the occurrence of vacuoles in the cytoplasm confirmed a toxic effect (Fig. 7C). Both the agarose composite (C2-A0.5-G1 and C2-A1-G1) treated groups were observed to have level 0 cytotoxicity confirming their biocompatibility.
Fig. 7Indirect cytotoxicity test: [A] Viability of NCTC clone 929 cells and [B] HADF cells treated with different concentrations of C2-A0.5-G1 and C2-A1-G1 extracts; Microscopic images of [C] NCTC clone 929 cells and [D] HADF cells treated with extracts in 10X (scale bar - 50 µm) and 20X (scale bar - 20 µm) magnifications. (ns> 0.05).
In addition, the indirect cytotoxicity analysis of human adult dermal fibroblast cells also confirmed the biocompatibility of C-A-G composite hydrogels. Cells treated with 100, 75, 50 and 25% extracts of the 0.5% agarose composite (C2-A0.5-G1) hydrogels showed higher cell viability of 105 ± 7%, 100 ± 8%, 95 ± 6% and 95 ± 6% (Fig. 7B). After 24 h of treatment 1% agarose composite hydrogel extracts (C2-A1-G1) of 25, 50, 75, and 100% showed viability of 102 ± 3%, 98 ± 5%, 89 ± 10% and 95 ± 7%, respectively. From the results, HADF cells treated with extracts of C2-A0.5-G1 and C2-A1-G1 hydrogels showed > 70% viability confirming the cytocompatibility of developed hydrogels.
In addition, the morphological evaluation of HADF cells treated with 100% extracts and control (culture medium) showed similar native morphology. Like the culture medium treated group, 100% extract-treated cells exhibited fully extended spindle-shaped cell morphology with improved cell-cell communication (Fig. 7D). In contrast, cells treated with phenol showed morphological changes such as formation of vacuoles and debris accumulation, which indicate the toxicity of phenol. The indirect and direct cytotoxicity evaluation confirmed the non-toxic effects of optimized triad bioink composition for biofabrication of tissues & organs. In a recent study, Boonlai et al., performed an indirect cytotoxicity study to evaluate thermoresponsive Pluronics F127/methylcellulose (PF/MC) hydrogel bioinks. Cytocompatibility was measured to be 100–102% for treated groups including culture media, 18PF/MC extract and 20PF/MC extracts. Lower cell viability of 9-10% was observed in the ethanol-treated group [
]. Similar to published reports, it may be concluded that C2-A0.5-G1 and C2-A1-G1 are non-toxic and hence suitable for tissue engineering applications.
Fig. 8 shows the live/dead staining of cultured NCTC clone 929 and HADF cells treated with cell culture medium (negative control), liquified phenol (positive control) and 100% agarose composite extract. Spindle-shaped fibroblast cells with higher extensions similar to the control group was observed in the 100% extract-treated group. In both the C2-A0.5-G1 and C2-A1-G1 groups, few dead cells were observed in NCTC clone 929 and HADF cells. These results confirmed the cytocompatibility of C2-A0.5-G1 and C2-A1-G1 composite hydrogels. The phenol-treated group was found to have only dead cells representing its toxic effect.
Fig. 8NCTC clone 929 and HADF cells stained with calcein-AM (green) and EtBr (red) after treating with 100% extracts (10 X magnification, scale bar: 50 µm)
Direct contact cytotoxicity study evaluates material cytocompatibility when directly contacting the live cells. Similar to the extract-treated (indirect) cytotoxicity evaluation, more than 30% cell death is considered to be a toxic material as per the ISO. The direct contact cytotoxicity test was performed for both C2-A0.5-G1 and C2-A1-G1 composites, which showed higher cell viability of 88 ± 5% and 100 ± 5%, respectively, for NCTC clone 929 cells (Fig. 9). Further microscopic evaluation of cultured cells after direct contact with C2-A0.5-G1 and C2-A1-G1 composite hydrogels confirmed the absence of cell morphological changes. Compared to the non-treated control group, the hydrogel-treated group showed similar cell extensions in the inner and outer hydrogel contact regions indicating biocompatibility.
Fig. 9Direct cytotoxicity test of NCTC clone 929 cells: [A] Viability, [B] Microscopic images of cells treated with hydrogels (scale bar - 50 µm). *p < 0.05
Fig. 10 shows the direct cytotoxicity evaluation for the composite hydrogels when in direct contact with HADF cells. HADF cells demonstrated high cell viability of 87 ± 8% and 81 ± 5% upon 24 h of culture in direct contact with C2-A0.5-G1 and C2-A1-G1 hydrogels, respectively. Further morphological evaluation of cells also confirmed the presence of native extended cell morphology with better cell-cell interactions. The present study confirmed the non-toxic nature of the developed hydrogels by showing more than 70% cell viability. Similarly, for cytotoxic evaluation of chondroitin/ dextran surface incorporated nanocellulose-alginate (NC-Alg-CS or NC-Alg-DS) based polysaccharide bioink composition for cartilage regeneration was tested for toxicity with indirect and direct contact cytotoxicity assay using cultured NCTC clone 929 fibroblast cells. This study results confirmed the non-toxic effects of the hydrogels by showing higher cell viability of more than 70% [
]. Likewise, Lafuente-Merchan et al., performed bioink cytotoxicity evaluation using indirect and direct assay methods as per ISO guidelines. Briefly, for indirect assay, extracts of fabricated nanocelluose-alginate (NC-Alg) bioink with and without hyaluronic acid (HA) composition showed 80–140% cell viability. Printed hydrogels were directly treated with the NCTC clone 929 fibroblast cells for direct contact assay test and found to have cell viability of 88.5 ± 33% for NC-Alg bioink and 74.1 ± 32% for NC-Alg-HA bioink. In both assays, more than 70% cell viability in the bioink treated groups confirmed the absence of potentially harmful effects on cell viability [
Development, characterization and sterilisation of Nanocellulose-alginate-(hyaluronic acid)- bioinks and 3D bioprinted scaffolds for tissue engineering.
Fig. 10Direct cytotoxicity test of Human Adult Dermal Fibroblast (HDAF) cells: [A] Viability, [B] Microscopic images of cells treated with hydrogels (scale bar - 50 µm). (ns> 0.05).
3.7 Complex shape printability of carboxymethyl cellulose -agarose-gelatin hydrogel
The availability of cost-effective, natural and biomimetic bioinks is one of the key challenges in the field of bioprinting and tissue engineering. In vitro biocompatibility studies performed on 0.5% agarose and 1% agarose composite confirmed their safety to utilize for biofabrication applications. In specific, C2-A1-G1 bioink developed in the present study was successfully used to print large and complex shapes due to its better thermoresponsive behavior. This composite hydrogel showed a higher storage modulus with a lower degradation rate, which is suitable to produce stable volumetric structures. As mentioned earlier, agarose and gelatin are thermoresponsive biopolymers that depend largely on the concentration and molecular weight of the polymers. carboxymethyl cellulose is a cellulose derivative widely used as a thickener in the food industry and it has been used recently for various biomedical applications such as skin regeneration, eye drops preparation, drug delivery, and bioprinting of various tissues [
]. In the present study, the bioink was maintained at 37 °C for 10 min to keep the bioink in pre-gel condition and further extruded at a pressure of 0.4 bars using a 0.4 mm NID (print bed temperature -10 °C). Other printing parameters include the printing speed of 15 mm/s, pre-flow and post-flow of 0.5 s. Complex models were created using TinkerCAD software and printed using the optimized printing parameters. All the 3D models were sliced using Bioplotter RP software with a layer thickness of 0.35 mm to print complex models. Models such as a hole-in box (1 cm thick– 28 layers), clover (2 mm thick – 6 layers), fan (2 mm thick– 5 layers) and letter “O” (2.5 mm thick– 8 layers) were successfully printed to confirm the printability of large volume structures. Further, printing groove structures (7 mm thick– 20 layers) and stepwise patterns (1.25 cm thick – 35 layers) confirmed the printability of the developed bioink to create complex shapes with uneven surfaces as in native human scale tissues. Based on the printability results, the developed bioinks have the capability to fabricate different complex shapes and large-sized models with better resolution (Fig. 11).
Fig. 11Bioprinting of various complex models (A) Fan, (B) hole in box, (C) Clover, (D) Stepwise patterns, (E) Letter “O” and (F) Grooves using C2-A1-G1 bioink (scale bar -1 cm).
In the present study, a semi-IPN-based triad of C-A-G composite hydrogel was successfully fabricated and evaluated their applications in biofabrication of volumetric constructs. Thermoresponsiveness of agarose and gelatin was well utilized for instantaneous crosslinking and smooth extrusion of bioink through the printing nozzle which enables less shear stress during the printing process. The gelation test confirmed the rapid and strong gel formation in all C-A-G bioink compositions. Further, the morphological and rheological analysis of 1, 3 and 5% gelatin containing C2-A0.5 and C1-A1 confirmed the variation in the porous architecture and hydrogel stability due to the difference in C-G entanglements. From these triads of various bioink compositions, the optimized C2-A1-G1 composite bioink showed improved thermal stability, higher porosity and swellability. These exemplary properties of the composite hydrogels are due to the formation of stable agarose crosslinking networks with carboxymethyl cellulose-gelatin entanglements. In vitro studies performed by indirect cytotoxicity and direct contact cytotoxicity assay confirmed biocompatibility with more than 70% cell viability for the composite hydrogels. Thus, this triad combination could be a suitable bioink candidate for the fabrication of volumetric soft tissues constructs.
CRediT authorship contribution statement
Muthu Parkkavi Sekar: Methodology, Validation, Data curation, Formal analysis, Writing – original draft. Harshavardhan Budharaju: Methodology, Data curation, Writing – original draft. Swaminathan Sethuraman: Data curation, Formal analysis, Writing – review & editing. Dhakshinamoorthy Sundaramurthi: Conceptualization, Methodology, Formal analysis, Supervision, Funding acquisition, Writing – review & editing.
Declaration of Competing Interest
The authors declare that there is no conflict of interest.
Acknowledgements
The authors wish to acknowledge Nano Mission, Department of Science and Technology (DST) (SR/NM/TP-83/2016 (G)), and Prof. T. R. Rajagopalan R & D Cell of SASTRA Deemed University for financial and infrastructural support. We also wish to acknowledge ATGC grant, Department of Biotechnology (DBT) (BT/ATGC/127/SP41147/2021), Adhoc funding, Indian Council of Medical Research (ICMR) (17 × 3/Adhoc/23/2022-ITR) and DST SERB CRG (Exponential Technologies) grant (CRG/2021/007847) for financial support. The authors are thankful to the Council of Scientific & Industrial Research (CSIR), Government of India for the senior research fellowship (09/1095(0051)/19-EMR-I and Indian Council of Medical Research (ICMR) for the senior research fellowship (3/1/1(4)/CVD/2020-NCD-1).
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Shapable bulk agarose–gelatine–hydroxyapatite–minocycline nanocomposite fabricated using a mineralizing system aided with electrophoresis for bone tissue regeneration.
Synthesis of cross-linked carboxymethyl cellulose and poly (2-acrylamido-2-methylpropane sulfonic acid) hydrogel for sustained drug release optimized by Box-Behnken Design.
Development, characterization and sterilisation of Nanocellulose-alginate-(hyaluronic acid)- bioinks and 3D bioprinted scaffolds for tissue engineering.
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