Polysaccharide-based Hybrid Self-healing Hydrogel Supports

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Polysaccharide-based Hybrid Self-healing Hydrogel Supports Paracrine Response of Mesenchymal Stem Cells Jijo Thomas, Anjana Sharma, vineeta Panwar, Vianni Chopra, and Deepa Ghosh ACS Appl. Bio Mater., Just Accepted Manuscript • DOI: 10.1021/acsabm.9b00074 • Publication Date (Web): 02 Apr 2019 Downloaded from http://pubs.acs.org on April 3, 2019

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ACS Applied Bio Materials

Polysaccharide-based Hybrid Self-healing Hydrogel Supports Paracrine Response of Mesenchymal Stem Cells Jijo Thomas, Anjana Sharma, Vineeta Panwar, Vianni Chopra, Deepa Ghosh*. Institute of Nano Science and Technology, Habitat Centre, Phase 10, Mohali 160062, Punjab, India. *Correspondence: Dr. Deepa Ghosh ([email protected], Tel.: 0172-221075)

Keywords: Injectable Hydrogel, Schiff-base, Mesenchymal stem cells, Paracrine response, Cell migration

Abstract The aim of stem cell therapy is to repair damaged tissues. Some of the challenges facing its success include cell retention and survival at the wound site. While retention of cells has been addressed by employing scaffolds, survival of transplanted cells in the repair tissue is however low. It is hypothesized that the observed regeneration is more a result of migration of tissue repairing cells from adjoining tissues in response to paracrine factors secreted by implanted cells, than by the implanted cells per se. In this study, we report the synthesis of a self-healing hybrid hydrogel that is injectable. The hybrid hydrogel was developed using the dynamic equilibrium of Schiff base linkage between the aldehyde groups on carboxymethyl cellulose dialdehyde (CMC-D) and amino groups on carboxymethyl chitosan (CMCh). Hydrogel stiffness and kinetics of gelation was observed to be modulated with different molecular weights of chitosan. In vitro studies demonstrated the cytocompatibility, hemocompatibility and biodegradability of the hydrogel. Chemotactic, proliferative and wound-healing response of cells to the paracrine factors secreted from MSC-hydrogel composite confirmed the ability of the hydrogel to support the paracrine response of stem cells. Our results suggest that the synthesized hydrogel-MSC composite could serve as a potential scaffold for studying the in vitro response of cells to the paracrine factors released by the encapsulated cells as well as a cell delivery vehicle for in vivo applications. Introduction Implanted mesenchymal stem cells (MSCs) are reported to improve tissue repair and regeneration of tissues such as the heart, cartilage, bone and skin.1 While the homing capability of MSCs has been exploited for its systemic application; evidence indicates that the majority of administered MSC get trapped in capillary networks and have a short life span2 leading to exploration of alternate approaches for delivering MSCs. Direct injection of MSCs at the repair site in gaining popularity for clinical applications. However, low retention and low engraftment of directly injected cells still present major hurdles for its successful clinical translation.3 To improve its retention and viability, MSCs have been implanted in association with biomaterials to treat damaged tissues such as skin4, heart5, cartilage6-7 etc.

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Advantages of injectable scaffolds over preformed scaffolds are: a. Ability to support minimally invasive delivery thus reducing surgery-associated risks; b. Fill irregular defects and adjust to the tissue topology; c. Provide easy access to deep tissues and d. Reduce treatment duration and its related costs. One potential attractive strategy for stem cell delivery is to suspend the stem cells in hydrogels which can be injected. Hydrogels, with high water content mimic the extracellular matrix (ECM) and have the potential to organize the cells into a three-dimensional architecture. The injectability of hydrogels further support facile delivery of cells via injection; provide homogeneous cell distribution within the desired tissue; and strong adhesion to the tissue after gelation, which increases the hydrogel-tissue interface, thereby improving the interaction of the hydrogel and host tissue.8 Treatment of damaged tissues with MSCs was earlier based on the concept of its ability to differentiate into several cell types9-10 and its fusion with host cells.11-12 With the limited survival and differentiation of MSC at the lesion site, it is now proposed that the therapeutic response is primarily due to its paracrine signalling. This paradigm is supported by in vitro and in vivo studies showing that several cell types respond to paracrine signalling from MSC resulting in their modulation of cellular responses, such as survival, migration, proliferation and differentiation13-14. Paracrine signalling is a form of communication between two different cells, where soluble factors released by a cell to its immediate environment results in a change in the behaviour of cells in its adjacent environment. We and others have confirmed that MSCs release several factors including migration factors15, proliferation factors16-17 etc. that results in the migration of cells from adjacent tissues. Hence, the focus of recent stem cell studies is on the secretome of stem cells and its effect on regeneration.18-19 To develop an injectable hydrogel for regenerating a damaged tissue with MSCs, it is imperative that the hydrogel in addition to its biocompatible and biodegradable properties should also support paracrineinduced cell response. Hydrogel formation can be accomplished through strategies including triggered gelation, or the use of shear-thinning hydrogels that are also rapidly self-healing. Triggered gelation can be induced in situ using biological stimuli, such as ion concentration20, light21, temperature22, pH23 etc. Selfhealing hydrogels are another class of injectable hydrogels that can be injected as a result of its shearthinning property. When subjected to shear stress these hydrogels undergo viscous flow and time-dependent recovery and reassembly of the hydrogel network upon relaxation. Self-healing hydrogels allow encapsulation of stem cells through weak dynamic interactions (e.g., hydrophobic interactions, hydrogen bonding, and electrostatic attractions) between the polymer chains before cell delivery.24 Self-healing hydrogels based on Schiff base are of great interest because gelation and healing readily occur in situ at physiological conditions without requiring external stimuli.25


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Hydrogel systems based on Schiff base crosslinking between amine and aldehyde groups have been reported for cell encapsulation.26-27 In all such studies at least one component is a polysaccharide derivative like alginate, dextran, chitosan etc. Reasons for such a preference would be due to the presence of a large number of crosslinking sites on these natural biopolymer chains which facilitate higher crosslinking density and greater hydrogel stability, and the ease of syntheses by oxidation of the sugar rings in polysaccharides. Very few studies have demonstrated the use of hydrogels to study paracrine response of cells. In these studies, the hydrogels were incorporated with stem cell-derived secretome and the response of cells like epithelial and endothelial cells to such hydrogels were assessed in vitro.28-30 The secretome released by MSCs is dependent on the diverse stimuli that the MSCs encounter at the site of implantation.31 To obtain the desirable MSC response, it is imperative that the implanted cells survive and secrete paracrine factors. With an aim to develop an injectable hydrogel that can support the viability and paracrine response of MSCs, we synthesized a series of hydrogels based on a reported combination32 by linking aldehyde groups on carboxymethyl cellulose dialdehyde (CMC-D) with the amino groups on carboxymethyl chitosan (CMCh). The properties of Schiff base hydrogels are reported to be primarily influenced by the number of aldehyde groups.27, 33 On the contrary, we demonstrate the impact of molecular weight of chitosan polymer on gel characteristics. To the best of our knowledge, we demonstrate for the first time the ability of the MSC-encapsulated hydrogel to support biological activity of cells in response to the paracrine factors released from the MSC-hydrogel composite.

Experimental Section Materials: Chitosan (Low, medium and High viscosity with degree of deactetylation >90%) was purchased from Mahtani chitosan Pvt Ltd, India and HiMedia Laboratories, India, Carboxymethyl cellulose was from Himedia Laboratories. Sodium metaperiodate, Hydroxylamine hydrochloride were from Merck, USA. DMEM, DMEM/F12, Trypsin, Penicillin-streptomycin-Amphotericin B, Vybrant® cell-labelling solution and all tissue culture plasticware were purchased from Thermo Fisher Scientific, USA. Cell counting kit8 (CCK-8) was purchased from Sigma. Synthesis of Hydrogel Synthesis of Carboxymethyl chitosan (CMCh): Synthesis of N,O-carboxymethyl chitosan (CMCh) was performed using a reported procedure with slight modifications.34 Viscosity of different grades of chitosan was verified using rheometer with plate and plate geometry. Chitosan (2g) of different viscosity grades [high (Mv~380kDa), medium (Mv~190kDa), low (Mv~86kDa)] was dispersed in 22 ml isopropyl alcohol (IPA), 50% Sodium hydroxide solution (12.5mol/ mol of chitosan) was added and allowed to alkalise for 2 h at 50oC. 2-chloroacetic acid dissolved in isopropyl alcohol (1:1, 5.5mol/mol of chitosan) was added



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dropwise to the chitosan dispersion under stirring at 50°C. The solvent ratio of IPA: water was maintained at 80:20. After 4 h the reaction was stopped by adding excess ethanol and washed several times with IPA to remove unreacted components. The precipitated material was dissolved in water and neutralised with dilute HCl. The insoluble fraction was removed by centrifugation and the supernatant was mixed with excess ethanol to precipitate out carboxymethyl chitosan. CMCh was purified by dialysis against ultrapure water for 3 days and lyophilized. The IR spectrum of CMCh was recorded with Cary 660 FTIR (Agilent technologies) in the range of 4000- 400 cm-1 using KBr pellet method. Synthesis of Carboxymethyl cellulose dialdehyde: The methodology followed in the synthesis of carboxymethyl cellulose dialdehyde (CMC-D) was as per reported procedure with few modifications.35 The synthesis parameters were selected to obtain a theoretical oxidation value of 45 % in order to obtain improved yield and stability, and to avoid unwanted oxidation products. Briefly, 2 g of carboxymethyl cellulose (CMC) was dissolved in 100 ml of ultrapure water and pH was adjusted to 3 with 0.1N H2SO4. Sodium metaperiodate 0.96 g (~0.45mol/mol of CMC) dissolved in 10 ml of water was added to the solution and allowed to react at 35°C for 3h. The reaction vessel was covered with aluminium foil to provide a completely dark environment. The reaction was stopped by adding ethylene glycol (equimolar to sodium metaperiodate) and the oxidised material was precipitated with absolute ethanol. The precipitated CMC-D was washed with water and ethanol several times to remove iodine compounds. CMC-D was purified by dialysis against ultrapure water for 3 days and lyophilised. KBr pellet method was employed to record the IR spectrum of CMC-D using Cary 660 FTIR (Agilent Technologies, USA) in the range of 4000- 400 cm- 1. The aldehyde content in CMC-D was determined by dissolving 0.5 g of CMC-D in 25 ml of water and pH was adjusted to 5. Hydroxylamine hydrochloride solution (25ml, 0.72M) was added to the CMC-D solution and allowed to react for 4 hours under stirring in a water bath maintained at 40°C. Released HCl was determined by titrating with 1M NaOH solution using methyl orange indicator and the volume (Vc) was recorded. CMC solution of same concentration was prepared and the titre volume was noted as Vb. Aldehyde content was calculated using the following equation (1).35

𝐴𝐶(%) =

𝑀𝑁𝑎𝑂𝐻 (𝑉𝐶 − 𝑉𝑏 ) × 100 𝑚/211

(1)

Where, MNaOH is 1 Mol/L, m is the dry weight of sample unit in g and 211 is the approximate molecular weight of repeating unit in CMC-D. The experiment was done in triplicate. In situ formation of self-healing hydrogel: Schiff base driven in situ gelation was employed for the hydrogel formation. Hydrogels were attempted using different concentrations of CMCh and CMC-D.


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Preliminary studies showed that the molecular weight of CMCh and the ratio of mixing CMCh and CMCD played an important role in hydrogel formation. To facilitate processing of the hydrogel, the concentration of CMCh had to be optimized across various viscosity grades. For comparing the properties of hydrogels prepared with chitosan of low, medium and high viscosity grades (depicted as CMChL, CMChM, CMChH) equal volumes of 2% solutions of both CMCh and CMC-D were mixed thoroughly to induce in situ gel formation. The as-prepared hydrogels were coded as CMChL-D, CMChM-D and CMChH-D respectively. Gelation time of the respective hydrogels was determined by the tube inversion method. After mixing the respective solutions, the tubes were inverted to observe gel formation and the time was recorded as gelation time. All studies were performed in triplicate.

Characterisation of hydrogel Rheology: Rheological properties of the hydrogels were studied with Modular Rheometer MCR302 (AntonPaar, Austria) using cone and plate geometry (CP50-2). Hydrogels were prepared by mixing equal volumes of CMCh and CMC-D solutions in a cylindrical mould with internal diameter of 16 mm and height of 5 mm and transferred to the plate and the spindle was lowered to the measurement position with a gap of 0.21mm. Stress sweeps were performed at a constant frequency of 1 Hz in oscillatory rotational mode starting from the amplitude of 0.1% (1 rad s-1@37°C) to find out the cross over point of storage (G') and loss (G'') moduli which signify the Gel-Sol transition of the hydrogel. In this study linear viscoelastic region (LVE) was determined for collecting subsequent data. Frequency sweeps were performed by changing the angular frequency from 0.1 to 100 rad/sec in oscillatory rotational mode at a constant strain (amplitude) corresponding to LVE region of the hydrogel at 37°C in the linear viscoelastic regime to determine values of the elastic (G′) and viscous (G″) modulus. Degree of cross-linking: Trinitrobenzenesulfonic acid (TNBS) assay36-37 was employed in determining the amine content and the degree of crosslinking in each gel formulation. Briefly, ̴ 5 mg of lyophilised hydrogel was weighed and added to 1 ml distilled water, followed by 0.05% 1 ml of TNBS solution and 1 ml of 4% of sodium bicarbonate solution. The sample was heated on water bath at 40ºC for 4 h. To the solution 9 ml of 6 N HCl, at 40ºC for added and after 4 h absorbance was determined at 334 nm. Glycine was used to make the standard curve. Blank experiment was performed by adding 6N HCl to the sample prior to TNBS solution. The degree of cross-linking was calculated using the equation (2): 𝐷𝑒𝑔𝑟𝑒𝑒 𝑜𝑓 𝑐𝑟𝑜𝑠𝑠-𝑙𝑖𝑛𝑘𝑖𝑛𝑔 =

Abs of cross-linke𝑑 𝑔𝑒𝑙 A𝑏𝑠 𝑜𝑓 𝑛𝑜𝑛-cross-linke𝑑 𝑔𝑒𝑙


× 100

(2)


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Self-healing and injectable property of hydrogel: Self- healing property of the hydrogels at macroscopic level was demonstrated by the ability of two hydrogel pieces placed in close proximity to join together.3839

In our study, hydrogel discs (5 mm diameter, 3 mm thickness) were stained either orange (by methylene

orange) or blue (by methylene blue) respectively. The discs were cut in 2 halves and two pieces of alternate coloured gels were placed in close proximity at room temperature (25°C) and at 37°C and the time taken for resealing was determined. The injectability of respective hydrogels was assessed by its ability to pass through an 18-gauge needle. Self-healing property of the hydrogel was further confirmed by its recovery after high shear load following oscillatory strain switching between low and high strain.40 In this test, individual hydrogels were exposed to a low strain of 1% at 1Hz, followed by high strain for 60 sec resulting in their deformation. The strain which corresponds to the cross over point of storage (G') and loss (G'') modulus of the respective hydrogels was noted. The strain observed with CMChH-D was 250% and with CMChM-D & CMChL-D the strain was 150%. The strain was then reduced to 1% for recovery of hydrogel for 120 sec. The cycle was repeated 3 times for each type of hydrogel. Morphology, Swelling ratio and in vitro degradation of hydrogels: The morphology analysis of the hydrogels were performed after freeze-drying under controlled environment. The hydrogels were frozen and freeze-dried using freeze-drier (FDUT-12003, Operon, Republic of Korea) for 24 h. The samples were sectioned, sputter coated with gold and visualised using JSM-IT300 scanning electron microscope (JEOL Ltd, Japan). The pore size distribution was assessed from the SEM images. Swelling studies with hydrogel was carried out in PBS (phosphate buffered saline) at 37oC. To 1G of hydrogel placed in a 35mm petri dish, 1 ml of PBS was added and incubated at 37oC. At different time points, PBS was removed from the petri dish, blotted with filter paper and weighed. The swelling ratio was calculated using the equation (3): w 𝑚𝑖 − 𝑚𝑜 Swelling ratio ( %) = × 100 w 𝑚𝑜

(3)

Where, mi and m0 is the initial and swollen weight of the hydrogel, respectively. Biodegradation of the hydrogel was determined in vitro. One gram of hydrogel was placed in 35mm petri dish and exposed to PBS for 24h to swell. Following swelling, the weight of each hydrogel (W0) was determined. The hydrogel was then incubated in PBS or lysozyme (0.15% in PBS) solution at 37°C. At the end of each set time point, the media was removed and excess solution was blotted with filter paper and the hydrogel was weighed. In vitro degradation was calculated using the equation (4):


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𝐺𝑒𝑙 𝑟𝑒𝑚𝑎𝑖𝑛𝑖𝑛𝑔 =

𝑊𝑖 × 100 𝑊0

(4)

Where W0 is the initial weight of hydrogel and WI is the weight of hydrogel at the set time point. Biocompatibility of Hydrogel Cytotoxicity of hydrogels: Cytocompatibility of the hydrogel was studied as per the method proposed in ISO 10993-5 for in vitro cytotoxicity assessment of medical devices. Test hydrogel was kept in contact with DMEM media for a period of 72h at 37°C. After incubation with the hydrogel, the contact media was filtered through 0.2m syringe filter and used for cytotoxicity assessment of the eluting substances. The experiment was carried out on fibroblastic cell line L929 cells. L929 cells were seeded (2x103 cells/well) in 96-well TC plates. After O/N incubation at 37oC in 5% CO2 incubator, the cells were exposed to the hydrogel extracts for 72h. Morphology of the cells were monitored and cell proliferation was assessed using 3-(4, 5-Dimethylthiazol-2-yl)-2, 5-Diphenyltetrazolium bromide (MTT) method. Viable cells were indirectly quantified by the purple colour formed as a result of the formazan crystals using spectrophotometric analysis. Hemocompatibility assay: To assess hemocompatibility of the hydrogel, in vitro hemolysis assay was performed using the indirect contact method on hydrogel extract as reported.41 To 2 gram of the hydrogel in a 35-mm petri dish, 2 ml of PBS was added and incubated for 72h at 37°C. The PBS in contact with hydrogel was used for hemolysis test. Goat blood (5 ml) was collected from a local butcher shop and centrifuged at 1500 rpm for 5 min to separate the RBCs. The isolated RBCs were suspended in PBS to form a 2% suspension. To 1.5 ml of 2% RBC suspension, 1.5 ml of the hydrogel extract was added. PBS was taken as a negative control and water as a positive control in place of media. The mixture was incubated for 3h at 37°C and centrifuged at 1500 rpm for 5 min. Absorbance of the supernatant was measured at 545 nm and percentage hemolysis was calculated from the equation (5) below. The results are expressed as mean ± SD from triplicate experiments. Percentage Hemolysis =

𝐴𝑏𝑠 𝑜𝑓 𝑠𝑎𝑚𝑝𝑙𝑒−𝐴𝑏𝑠 𝑜𝑓 𝑛𝑒𝑔𝑎𝑡𝑖𝑣𝑒 𝑐𝑝𝑜𝑜𝑛𝑡𝑟𝑜𝑙 𝐴𝑏𝑠 𝑜𝑓 𝑃𝑜𝑠𝑖𝑡𝑣𝑒 𝑐𝑜𝑛𝑡𝑟𝑜𝑙−𝐴𝑏𝑠 𝑜𝑓 𝑛𝑒𝑔𝑎𝑡𝑖𝑣𝑒 𝑐𝑜𝑛𝑡𝑟𝑜𝑙

× 100

(5)

Cell isolation and propagation To study the response of cells to the paracrine factors released from MSCs encapsulated in the hydrogel, we used fibroblasts and chondrocytes as model cells. MSCs were derived from umbilical cord. Human tissues such as umbilical cord, skin and cartilage were used to derive cells. The tissues were obtained from donors after receiving informed consent as per the protocol approved by the Institutional Ethics Committee of PGIMER (IEC-08/2017-658), Chandigarh, India.



Cord tissue processing for MSC isolation and propagation was carried out using our previously reported protocol16. In brief, the umbilical cord was decontaminated and the cells were isolated by explant culture in Dulbecco's modified Eagle's medium F-12 (DMEM/F12) (Invitrogen) supplemented with 10% foetal bovine serum (Invitrogen), 2 ng/mL bFGF (Invitrogen), and 1X antibiotic–antimycotic solution (Invitrogen). MSCs were used from passages 2-6 in the experiments. Isolated cells were immunophenotyped using the human MSC analysis kit (BD-Stemflow™, Becton-Dickinson) as per the manufacturer’s protocol. Data was acquired on a FACS Aria III instrument (Becton-Dickinson) using Cell Quest software. Dermal fibroblasts were isolated from dermis of discarded neonatal foreskin as per the published protocol.42 Briefly, neonatal foreskin was decontaminated and incubated with dispase (Sigma) to separate the epidermis from dermis. Fibroblasts were isolated from the dermis after digestion for 2h with 0.1% collagenase (Sigma). The isolated cells were propagated in DMEM media containing 10% FBS and antibiotics. Isolation of chondrocytes from discarded cartilage was carried out as per protocol.43 Briefly, the cartilage tissue was minced into small pieces and digested in PBS containing 0.1% Collagenase type II (Sigma) for 24h. The digested material was strained using a sieve to remove undigested tissue and cells were separated by centrifugation. The cells were propagated in Chondrocyte growth media (Lonza). Assessment of MSC proliferation in hydrogel Hydrogel was prepared using CMChL-D in 48-well plate by in situ gelation with the MSCs encapsulated in it. Briefly, CMChL solution (2%) containing suspended MSCs (2x105 cells/ml) was added to a 48-well plate and mixed with CMC-D solution to form a gel. Following gel formation, DMEM media with 10% serum was added to the wells and incubated in CO2 incubator with regular media change every alternate day. On the 1st and 5th day after seeding, cells were assessed for viability using FDA (fluorescein-diacetate) / dead cell using PI (propidium iodide). Cells in the hydrogel were visualized using a confocal laser scanning microscope (Zeiss LSM880 confocal microscope, Carl Zeiss) and the images captured using Z-stacking mode under 10x magnification. For quantitative analysis, MSCs (2.5x103 cells/well) were encapsulated in the hydrogel and seeded in a 96-well plate as mentioned above. As a control, similar number of MSCs were directly seeded in wells. At defined time points of 1, 3 and 5 days after seeding, proliferation was determined using CCK-8 kit (Sigma). At the end of each time point, CCK-8 reagent was added to each well and optical density of the media was determined at 460 nm using Infinite M Plex plate reader (Tecan). Each experiment was done in triplicate and repeated thrice.


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Assessment of hydrogel’s support to paracrine response of Mesenchymal Stem Cells Chemotactic response of cells to paracrine signalling by MSCs cultured in hydrogel was studied in vitro. MSCs were labelled with a green cell tracker (DiO, Vybrant® cell-labelling solution) and seeded in the hydrogel (CMChL) as reported above in a 48-well plate. Seeding density of MSCs was 2.5x105 cells/ml of hydrogel. After gelation, DMEM/F12 media containing 10% serum was added to the wells. After 48h, the media above the hydrogel was removed and overlaid with red cell tracker (Dil, Vybrant® cell-labelling solution)

labelled

fibroblasts/chondrocytes

suspended

in

media

containing

1%

serum.

Fibroblast/chondrocytes were similarly overlaid on plain hydrogel (without MSCs) which served as control. After 12h the migratory response of cells to the chemotactic factors released by the encapsulated MSCs in the hydrogel was studied using a confocal microscope (Zeiss LSM880 confocal microscope, Carl Zeiss) by taking slice pictures from top to bottom of the hydrogel in z-stack mode (slice thickness 10µm, z= ̴ 600µm) at 5X magnification. 3D picture of the slices were generated to observe the migratory pattern of the overlaid cells in to the hydrogels. To further confirm other biological effects of the paracrine factors released from MSC-encapsulated hydrogel, the media in contact with the MSC encapsulated hydrogel (Conditioned media) was tested for induction of fibroblast proliferation and wound healing in vitro. In the proliferation studies, dermal fibroblasts were seeded in 48-well plates. After 24 h, the media was replaced with control media (in contact with blank hydrogel) or conditioned media. Proliferation of the fibroblasts was assessed after 48h and 72h after treatment using MTT assay. Induction of wound healing by the released paracrine factors was tested indirectly using the scratch wound assay.44 Briefly, primary fibroblasts were seeded in 24-well plates and allowed to proliferate. On reaching confluence, the cells were wounded by creating a scratch. Cells were then treated with control media or conditioned media derived from the MSC-hydrogel composite. Wound healing was monitored by observing the migration of cells in the scratch area. Statistical Analysis: Each statistical data is derived from sample size of three and the data is presented as mean ± standard deviation (SD). GraphPad Prism 6 was used to analyse data. Statistically significant is declared if p ≤ 0.05 as determined by one-way or two-way ANOVA test keeping 95% confidence interval.

Results and discussion Synthesis of Hydrogel: Chitosan, a natural product with structural similarity to glycosaminoglycan has extensive scope in tissue engineering, but its insolubility in neutral pH has however limited its application. The time for gelation is important for injectable hydrogels because too fast gelation would induce hydrogel formation before its implantation and processing, whereas slow gelation would result in gel precursors diffusing away from the injection site and uneven distribution of cells.39 The concentration and molecular



weight of the components play a vital role in the rate of hydrogel formation.45-46 To determine the influence of molecular weight of chitosan on hydrogel formation, we chose low, medium and high molecular weight chitosan. Grafting carboxymethyl group on chitosan to form CMCh renders it soluble in water at neutral pH. The carboxymethyl groups were introduced into the chitosan, by reaction between the hydroxyl group and the amine group of chitosan with chloroacetic acid (Fig 1A). The viscosity of 1% solutions of the low, medium and high viscosity CMCh was 4, 15 and 110 cP (determined using rheometer, by plate and plate geometry) respectively. Formation of CMCh was confirmed from FTIR studies (Fig 1B). Introduction of the carboxyl groups on chitosan can be identified by the characteristic peak at 1604 cm-1 and a characteristic peak moderate band at 1411 cm-1 due to symmetric and asymmetric axial deformations of the carboxyl groups respectively.34 The degree of substitution of carboxymethyl group determined by the titration method on low, medium and high molecular weight grades of chitosan was 0.9±0.04, 0.87±0.06 and 0.83±0.05 respectively. Carboxymethyl cellulose dialdehyde was prepared by periodate oxidation of CMC to introduce aldehyde groups in the polymer backbone which is required in Schiff base formation with the primary amine present on chitosan (Fig 2A). The reaction time, pH and molar equivalence was selected to ensure better stability and yield with a significant level of oxidation. The yield of CMC-D was 65±3% and after oxidation the viscosity of the solution was decreased from 467 cP to 3cP (determined by plate and plate viscometer by taking 1% solution). Formation of aldehyde groups was confirmed by the carbonyl stretching band at 1730cm-1 and hemiacetal band at 883 cm-1 in the FTIR spectra.35 (Fig 2B). Aldehyde content was determined by the hydroxylamine hydrochloride titration method. It was found to be 35±3%, out of 45% theoretical degree of oxidation.

Figure 1: (A) Synthesis of CMCh and (B) IR spectra of CMCh


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Figure 2: Synthesis of CMC-D (A) and FTIR of CMC-D (B) The hydrogels were prepared through homogeneous mixing of equal volumes of 2% weight/vol of CMCD and CMCh solutions of various molecular weights (as shown in scheme 1) so that the final concentration of the individual components in the hydrogel would be 1%. On mixing the respective components, the amine groups on chitosan backbone were cross-linked with the aldehyde groups at the terminal of CMC-D to form reversible imine bonds at ambient temperature (Fig 3A). The dynamic hydrogel network was based on the equilibrium between Schiff base linkages and the disassociated aldehyde and amine groups. As the gel formation occurs almost instantaneously, these hydrogels can be categorized as in-situ forming hydrogels. The absence of aldehyde band (1730cm-1) in hydrogel spectrum confirmed the formation of imine bond (Fig 3B).

Scheme 1: Schematic representation of the preparation of hydrogel



Figure 3. A. Synthesis of hydrogels. B. IR spectra of hydrogel

Gel Characterization Gelling time: The time required for gel formation, measured by the vial inversion test47 was found to be ~15.6±1.7 s, 11±2 s and 3.3±0.5 s for CMChL-D, CMChM-D and CMChH-D respectively (Fig 4). Morozowich et al earlier reported the influence of aldehyde content on gelation time of Schiff base hydrogels; with higher aldehyde percentage showing faster crosslinking.48 In this study where we have used the same amount of aldehyde, a clear relationship is observed between the different molecular weights of chitosan and the rate of gel formation, indicating that molecular weight of chitosan can also influence the rate of gelation. In hydrogel formation, besides Schiff base reaction which is primarily known to be responsible for gel formation, other factors like physical cross-linking (chain entanglement due to interpenetrating polymer network) is also reported to control the rate of gelation.49-50 This could explain our observed faster rate of gelation seen with increasing molecular weights of chitosan.

25

*** G e la t io n T im e ( s )

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20

*

15

**

10

5

0 C M C h H- D

C M C h M -D

C M C h L -D

Figure 4: Gelation time of the hydrogel. Graph represents data from an average of 3 experiments. *** represents P-value ≤0.001; ** P-value ≤ 0.01; and * P-value ≤0.05 respectively (One-way ANOVA, Tukey's multiple comparisons test).


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Rheology and mechanical properties: The mechanical properties of an injectable hydrogel play a critical role in its successful clinical application. The hydrogel should be able to convert to a low-viscosity liquid during its application, and reform after implantation. In addition, it should possess suitable mechanical strength to withstand biomechanical loading and provide temporary support for the cells. Amplitude sweep test was performed to study the physical nature of the gel. Amplitude sweep data provides the elastic modulus (G′) and viscous modulus (G′′) at different shear strain (%). Shear strain was varied from 0.01%–100% to determine the Linear Viscoelastic Region (LVER), wherein the gel can elastically strain and return to its original state once the strain is removed. All further rheological analysis was performed in this LVER. Elastic modulus indicates the solid component and viscous modulus represents the liquid component of a gel. When the solid component is higher than the liquid component, it indicates formation of a gel. As seen in Fig 5A, all the compositions show higher elastic modulus than the viscous modulus, thereby confirming its gel property. The elastic modulus amongst the composites were in the order of CMChH-D>CMChM-D>CMChL-D indicating the elastic modulus is dependent on the molecular weight of chitosan. Shear-thinning behaviour of the gel was observed only when the strain was over 100%, at which point the network of the hydrogel is broken. The cross over point of storage modulus (G') and loss modulus (G'') confirms the gel to sol transition. From the inset-data (Fig 5A), it was observed that CMChLD requires a lower strain for sol-gel transition thereby favouring its injectable property. Frequency sweep data which determines the elastic modulus with respect to time is shown in Fig 5B. With increasing angular frequency, storage moduli G’ showed little change and loss moduli varied gently with no sign of breakage as far as the measured angular frequency range was concerned confirming that the hydrogels were cross-linked and mechanically robust. The G’ values obtained with different molecular weights of chitosan were in the order CMChH-D>CMChM-D>CMChL-D. The storage moduli of the composite hydrogels were in the range of 200-600Pa (Storage moduli of CMCh L-D, CMCh M-D and CMCh H-D hydrogels was 221±13 Pa, 339±38 Pa and 528±38 Pa respectively) with the moduli increasing with increasing molecular weight of chitosan (Fig 5C). The crosslinking density is mainly responsible for storage modulus, and the storage modulus generally increases with increasing crosslinking density. The degree of cross-linking was found to be CMCh L-D, CMCh M-D and CMCh H-D hydrogels was 69±2.6%, 76±2.3% and 86±2.9% respectively. We have observed the degree of crosslinking to increase with the molecular weight of chitosan (Fig 5D). The observed increase in cross-linking with molecular weight might be due to further entanglement of the polymeric chains leading to increase in the number of cross-link points thereby facilitating greater number of imine linkages. The chemical cross-linking observed in the hydrogel



might be further reinforced by physical cross-linking of the polymers through hydrogel bonding and/or chain entanglement.39 While the mechanical strength and viscoelastic property of the hydrogel plays an important role in maintaining the gel’s integrity, it also plays a crucial role on cells by effecting its morphology, differentiation, motility and viability.51-54 Organs in mammals have elastic moduli ranging between 10010,000 Pa, depending on the type of tissues.55 Cells like neurons, epithelial cells etc. are reported to survive on soft materials having a shear modulus of 100 Pa.56-57 Stiffness of hydrogels is reported to effect stem cell proliferation58 and differentiation.59-60 Wingate et al have reported that endothelial cells cultured in the conditioned media from MSCs differentiated in the soft matrix and formed capillary-like structures, indicating the importance of mechanical properties of the substrate for cell survival and eliciting its biochemical response.60 The mechanical characteristics observed with our synthesized hydrogels indicate their suitability for further exploration of MSC applications.

Figure 5: Rheological properties of the gels (A) Amplitude sweep, showing the elastic modulus (solid lines)

and viscous modulus (dotted lines) at different shear strain. Inset shows gel to sol transition. (B) Frequency sweep measurements of the injectable hydrogels demonstrating their viscoelastic behaviour (G′ > G″). Solid


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lines show the elastic modulus and dotted lines show the viscous modulus of respective gels. (C) Storage modulus of the respective hydrogels. (D). Degree of cross-linking observed in hydrogels measured using TNBS assay. Data represents an average of 3 experiments. **** represents P-value ≤0.0001; *** P-value ≤0.001 ** P-value ≤0.01; and * P-value ≤0.05 respectively (One-way ANOVA, Tukey's multiple comparisons test).

Self-healing property: Cells in regenerating tissues are constantly interacting with and remodelling their surrounding matrix to support its activities like migration, proliferation, and differentiation.61 To overcome the deterioration in mechanical properties that exist in conventional, degradable hydrogels that are permanently crosslinked; new approaches are required that enable hydrogels to support normal cellular functions without requiring irreversible hydrogel degradation.62 Such hydrogels, termed as “adaptable hydrogels” have polymer networks with adaptable linkages that can be broken and re-formed reversibly without external triggers. To check the adaptable nature of our hydrogels, we performed macroscopic testing of self-healing behaviour of the hydrogels as shown in Fig 6. The synthesized hydrogels were coloured with 2 different coloured dyes and cast as discs (Fig 6A-a). The discs were cut transversely and different coloured sections were kept in close proximity (Fig 6A-b) to assess the rate of self-healing by observing the time taken for boundaries between the respective hydrogels to disappear. In the absence of external intervention, the time taken for sealing was 6, 8 and 15 min for CMChL-D, CMChM-D & CMChHD respectively (Fig 6A-c). The healed hydrogels could be suspended in air and did not separate at the boundary after application of an external force, indicating its self-healing ability. The self-healing process which occurs as a result of re-crosslinking of polymer chains is reported to be regulated by the amount of amine groups present at the interface.63 Our results indicate that self-healing rate of our hydrogels is inversely proportional to the molecular weight of chitosan. This could be due to the easy availability of the amine groups in compositions with lower molecular weight of chitosan as a result of lower hindrance to its dynamic mobility. To study the dynamic nature and the injectability of the hydrogels, individual hydrogels were loaded into syringe and extruded through 18-gauge needle (Fig 6d). As shown in Figure 6(e), all the three hydrogels could be extruded suggesting their injectability. The ease of injectability was in the order of CMChLD>CMChM-D > CMChH-D which is supported with the rheological data wherein the gel-sol transition was observed to be lowest in CMChL-D gel. After extrusion, the hydrogel strips healed to reform gels that could be held by forceps (Fig 6f-g). These results further confirmed the self-healing capability of the hydrogels. On compression of the hydrogels in the syringe, the pressure applied caused dynamic uncoupling of Schiffbase linkages leading to its deformation and hence the gel passed through the thin needle. As the rate of self-healing to re-form an integrated gel is dynamic and depends on the rate of recoupling of the reversible



linkages, gels prepared with increasing molecular weights of chitosan took a longer time to re-heal. Although, the hydrogels could heal itself without additional stimuli, the rate of self-healing was however observed to improve when the process was tested at 37°C as compared to RT (25°C) (Fig 6B). This could be due to improved facilitation of diffusion of the amine and the aldehyde active sites at higher temperature.

Figure 6: Self-healing behaviour of the injectable hydrogel. CMChL-D, CMChM-D & CMChH-D hydrogels are represented as L, M and H respectively. (A) Individual hydrogels were prepared in a mould in the form of discs (a). While one disc from each hydrogel is dyed orange, the other is dyed blue to easily distinguish the interface. (b) The discs were cut into two halves and were placed in close proximity to different colour gels at room temperature (25oC). (c) The hydrogels self-healed in 6-15 min as demonstrated by lifting the discs using forceps. (d) Shows the hydrogel being extruded through a syringe. (e) The extruded hydrogels and (f) the reformed gel being lifted using forceps (g). (B) Time required for self-healing of the discs when observed at RT (25oC) and at 37oC. **represents P-value ≤0.01; and * P-value ≤0.05 respectively (Two-way ANOVA, Bonferroni's multiple comparisons test). Each data represents an average of 3 experiments.

Rheological recovery tests were performed to evaluate elastic response and self-healing performance of hydrogel by exposing the hydrogel to a high deforming shear and then allowed to recover. The shear employed for CMChL-D, CMChM-D was 150% and for CMChH-D it was 250% respectively at 1Hz which is the gel-sol transition strain for the corresponding gel. Changes in the strain-dependent modulus was observed between the three hydrogels. Changes in the damage-healing (sol-gel transition) of


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the respective hydrogels, which was obtained by application of a large amplitude oscillatory strain and recovery at a low strain (strain=1%, frequency=1.0 Hz), is shown in (Fig 7). The G' value of the CMChHD hydrogel decreased from ~600 Pa to ~10 Pa when the strain was increased from 1 to 250% and recovered to ~500 Pa when the strain was decreased to 1% (Fig 7A). Similarly, on providing a strain of 150%, the self-healing process was similarly observed in the CMChM-D hydrogels, where the G' decreased from ~300 to ~50 Pa (Fig 7B) and recovered while for CMChL-D, it changed from ~200 Pa to ~50 Pa following decrease of strain to 1% & then recovered to ~200 Pa (Fig 7C). The quick recovery of the hydrogels inner network confirmed the self-healing capability of the dynamic hydrogels.

Figure 7: The recovery of hydrogel after high shear load demonstrated by the continuous step strain (A) CMChH-D (1% strain→250% strain→1% strain); (B) CMChM-D (1% strain→250% strain→1% strain) and (C) CMChL-D hydrogels respectively. Morphology, swelling and degradation properties of hydrogels: The influence of structural properties of hydrogels on cellular responses such as growth, migration, and biosynthetic activity has been reported.58, 64

An ideal injectable hydrogel should possess pores in the range of 50-300µm, and a high degree of

interconnectivity to facilitate transport of nutrients and oxygen, as well as cell adhesion and migration.65 Using SEM we studied the interior morphology and pore size distribution in hydrogels (Fig 8A). The



morphology suggests that the interior structure of the hydrogel is dependent on the degree of cross-linking. The higher the degree of cross-linking, the smaller the pore size as shown in Fig 8A. Lyophilized CMChLD hydrogel shows network with open and large pores with pore size range of 80-120 m (Fig 8A-a inset); followed by CMChM-D with 70-80 m (Fig 8A-b inset) and CMChH-D with 40-80m (Fig 8A-c inset) respectively. The network pore size is observed to corroborate with our crosslinking density. The presence of large continuous pores in the three-dimensional structure of CMChL-D hydrogels might provide adequate void space for cell proliferation as well as nutrient transportation. The swelling property of hydrogel is crucial for substance exchange when they are used as an injectable scaffold for cell applications. We evaluated the degree of swelling by immersing the hydrogels in PBS at 37°C. As seen in Fig 8B, the hydrogels underwent gradual swelling and reached a plateau in 10h (600 min). The maximum degree of swelling of CMChH-D, CMChM-D and CMChL-D hydrogels was 5%, 18% and 28%, respectively (Fig 8B). A similar observation was seen with the swelling rate. Both these phenomena may be attributed to the hydrogel crosslinking density, as degree of swelling is inversely proportional to the cross-linking density. It is recommended that after its implantation in the damaged tissues, the hydrogels should degrade with time and accommodate the regenerating tissue.66 Hydrolytic and enzymatic degradation tests were performed in vitro on CMChH-D, CMChM-D and CMChL-D hydrogels. All the three hydrogels demonstrated degradation with time (Fig8C). While the rate of degradation under each condition was in the order CMChL-D >CMChM-D > CMChH, the overall degradation rate was higher with lysozyme treatment (Fig 8Cii). The increased degradation rate on lysozyme treatment is due to enzymatic hydrolysis of the glycosidic bonds of acetylated residues present in the chitosan.67 The tunable degradation pattern of our hydrogels can thus be employed in different tissues with different cell and ECM turnover rates.


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Figure 8: Morphology, swelling and degradation properties of hydrogels A. SEM images of lyophilised hydrogels (a) CMChH-D, (b) CMChM-D, (c) CMChL-D. Scale bar represents 100µm. (B) Swelling characteristics of the hydrogels. (C) Degradation of hydrogels in the presence of (i) PBS and (ii) 0.15% lysozyme. Each data represents an average of 3 experiments.

Biocompatibility Cytotoxicity: The cytotoxicity of the leachable and degradation products of the hydrogel was investigated according to ISO 10993 using L929, a mouse fibroblast cell line. Cell images taken with an inverted microscope revealed nontoxic feature of the hydrogels (Figure 9A-a). The treated cells retained the spindle morphology of the cells, indicating the hydrogel’s cytocompatibility. In addition, quantitative analysis was performed based on MTT assay and the results confirmed that the hydrogel is nontoxic. Viability of L929 cells cultured on tissue culture plate served as control. The absence of cell death indicates that the degradation products of hydrogels are non-toxic to cells (Fig 9A-b)). The hemo-compatibility of the hydrogels was tested with blood using the classical hemolysis assay (Fig9B). RBCs treated with water which served as positive control shows 100% lysis, however no significant lysis was observed in RBCs treated with hydrogel extracts (Figure 9B) indicating hemo-compatibility of the hydrogels.



Figure 9: Biocompatibility of hydrogels. (A) Cytotoxicity. L929 cells were treated with media in contact with the respective hydrogels for 72 h. At the end of 24, 48 and 72 h the cells were visualized under a microscope. Scale:20 µm. (a) and the cells quantified using MTT assay (b). Data represents mean + SD from 3 experiments carried out in triplicate. ns indicates no significant change (One-way ANOVA, Tukey's multiple comparisons test). (B) Hemocompatibility of the hydrogels. MSC viability and proliferation in hydrogel Based on physical attributes of the hydrogel, including its ease of application, degree of swelling, internal morphology etc. which are favourable for cell growth and implantation, we selected CMChL-D for further evaluation for cell growth. MSCs were isolated from the umbilical cord and characterized to verify the identity of the cells (Figure S1). MSCs encapsulated in CMChL-D hydrogel and supplied with media. On the day of evaluation, the hydrogels were treated with FDA/PI to detect dead cells (stained red). On visualization of the cells under confocal microscope; live cells appeared as green fluorescent cells, while dead cells appeared as red cells as shown in Fig 10A. The cells were uniformly distributed in the hydrogels and exhibited good viability. Viability of the cells was evaluated by comparing the ratio of green to red cells. A higher ratio indicated that the hydrogel supported cell viability.68 Cell proliferation in hydrogel was assessed by increase in the number of green fluorescent cells on day 5 (Fig b & d) as compared to day 1 (Fig 10 a & c). Cell proliferation was further confirmed quantitatively using CCK-8 assay (Fig 10 B). At every time point tested, no significant difference in absorption was observed between cells cultured on regular cell culture dish or in the hydrogel. In addition, a consistent increase in absorbance with time was observed, indicating that the hydrogels supported cell proliferation. The proliferation data indicates that the hydrogel provides a favourable environment for the cells by supporting efficient nutrient/metabolic waste transportation through excellent pore connectivity.


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Figure 10: Effect of CMChL-D hydrogel on MSC proliferation. A: MSCs were encapsulated in CMChLD hydrogel and its viability was assessed by Live/Dead staining on days 1 and 5 after encapsulation. Representative confocal images of live (green)/ dead (red) cells viewing from top (maximum intensity projection). (a&c) and in 3D (b&d). Scale:100 μm. B. Cell proliferation of MSCs assessed using CCK-8 assay on days 1, 3 and 5 after encapsulation. The absorbance was compared with the absorbance obtained from MSCs cultured on TC dishes.

****

indicate P-value

Polysaccharide-based Hybrid Self-healing Hydrogel Supports

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