Injectable Disc-Derived ECM Hydrogel Functionalised with Chondroitin Sulfate for Intervertebral Disc Regeneration

Low back pain resulting from intervertebral disc (IVD) degeneration is a significant socioeconomic burden. The main effect of the degeneration process involves the alteration of the nucleus pulposus (NP) via cell-mediated enzymatic breakdown of key extracellular matrix (ECM) components. Thus, the development of injectable and biomimetic biomaterials that can instruct the regenerative cell component to produce tissue-specific ECM is pivotal for IVD repair. Chondroitin sulfate (CS) and type II collagen are the primary components of NP tissue and together create the ideal environment for cells to deposit de-novo matrix. Given their high matrix synthesis capacity potential post-expansion, nasal chondrocytes (NC) have been proposed as a potential cell source to promote NP repair. The overall goal of this study was to assess the effects of CS incorporation into disc derived self-assembled ECM hydrogels on the matrix deposition of NCs. Results showed an increased sGAG production with higher amounts of CS in the gel composition and that its presence was found to be critical for the synthesis of collagen type II. Taken together, our results demonstrate how the inclusion of CS into the composition of the material aids the preservation of a rounded cell morphology for NCs in 3D culture and enhances their ability to synthesise NP-like matrix.


Introduction
The intervertebral disc (IVD) is a highly hydrated fibro-cartilaginous tissue critical for the normal functioning of the spine. There is still no accepted consensus with respect to the biochemical, mechanical and nutritional mechanisms that triggers degeneration of the IVD.
However, it is clear that one of the effects of this process is the cell-mediated enzymatic breakdown of key extracellular matrix (ECM) components [1] and that it prevalently occurs in the nucleus pulposus (NP) of the disc [2][3][4]. The healthy NP consists of a network of highly hydrated sulfated glycosamminoglycans (sGAG) embedded in a loose type 2 collagen structure [5,6]. Chondroitin sulfate (CS) is the most abundant sGAG and is responsible for the high levels of hydration of the tissue and resulting biomechanical properties to withstand compressive loads [7]. During disc degeneration (DD) the sGAG content of the NP decreases resulting in reduced water content, and concomitant functional capacity [2,8].
Injectable biomaterials have been widely investigated in an attempt to restore or recreate the structural and biochemical composition of a damaged tissue, while at the same time having the advantage of being minimally invasive. Although a wide range of natural and synthetic polymers have been employed in tissue engineering, materials derived from the decellularisation of tissue specific ECM have received considerable attention in the last decade due to their high biocompatibility, excellent tissue integration, and intrinsic bio-inductive factors that can induce differentiation or promote a specific phenotype [9][10][11][12][13][14][15][16][17][18]. However, a major challenge of the decellularisation process is preserving the biochemical composition of the native tissue. During a standard decellularisation procedure native tissues are generally immersed into a series of water-based solutions. Due to their polar nature, GAGs are characterised by high water solubility and tend to disperse into the decellularisation solutions, compromising the integrity of the native tissue composition. Therefore, in order to retain the biomimetic composition of the material, it is necessary to reintroduce the fraction of GAGs lost during the processing of the ECM, and more specifically CS. The presence of CS has been shown to have a positive effect in increasing matrix production, providing chondro-protection and anti-inflammatory properties [19][20][21][22]. Nevertheless, it is widely known that the presence of GAGs influences collagen fibril formation, with fibril networks forming less readily in a carbohydrate rich environment resulting in hydrogels with diminished integrity and poor mechanical properties [23][24][25][26]. Previous work has shown that carboxyl activated or functionalised CS can be combined with collagen to form self-crosslinkable hydrogel systems [27] overcoming the challenge of producing ECM-CS gels with poor mechanical properties and stability.
In the research of an appropriate cell source for the regeneration of the disc a number of important factors, such as potential immune rejection, technical and regulatory issues of utilising allogeneic donor cells, make the use of a patient's own cellular material an obvious choice [28]. Among the alternative cell sources discussed and analysed in literature, nasal septal cartilage may provide for an attractive tissue source. This tissue is easily accessible to harvest under local anaesthetic and it has been demonstrated that nasal chondrocytes (NC) provide for good cell yields and a high chondrogenic potential even post expansion [28][29][30].
Recently, NCs have shown to represent a promising autologous cell source for articular cartilage repair in several animal studies [31,32] and may offer similar potential in IVD repair strategies [33].
The overall goal of this study was to develop an injectable self-assembled ECM hydrogel derived from native disc tissue and assess the effects of functionalised CS incorporation on the rheological properties and matrix forming capacity of nasal chondrocytes without growth factor supplementation.

ECM extraction, decellularisation and characterisation
Nucleus pulpous tissue was harvested from 18 month old bovine coccygeal discs under aseptic conditions. Discs were minced and treated with 20 ml/g of 200 mM sodium hydroxide (NaOH) at 4°C under constant rotation for 24h. NaOH treated fragments were then washed in diH2O, freeze-dried and cryomilled (Spex SamplePrep 6770 Freezer/Mill) to obtain a fine powder. The decellularisation process was performed by adding 1 g of ECM powder to 50 ml of 50 mM Tris Buffer, 1 mM MgCl2 solution supplemented with 2 U/ml of Benzonase® Nuclease (Merck Millipore, Ireland) and agitated at 37°C for 3 cycles of 1 hour each. Decellularisation solutions were changed at the end of each cycle. DNA content was determined using the quant-iT PicoGreen dsDNA Kit (Invitrogen) and base pair length was evaluated using agarose gel electrophoresis. Briefly, digests of decellularised ECM and native tissue as comparison were digested and purified using a Qiagen DNeasy® Blood and tissue extraction kit according to the manufacturer's instructions. Purified DNA samples were loaded into wells of 1% w/v agarose gel made with Tris Base-EDTA buffer and containing ethidium bromide. Gels were run at 80V for 60 minutes and imaged with a UV filter to detect DNA bands.

Functionalisation of chondroitin sulfate
Functionalisation of chondroitin sulfate (CS) was performed following a method previously described [9]. CS was dissolved in ultrapure water (UPW) containing 200 mM Nhydroxysuccinimide (NHS) and 200 mM 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC) at a concentration of 50 mg/ml (molar ratio EDC:NHS:CS groups = 2.5:1.9:1.0). The pH of the solution was maintained at 4.75 and the reaction was allowed to proceed for 2 hours at room temperature and terminated by adjusting the pH of the solution to 7.4. The product of the reaction was dialysed against diH2O at 4°C (MWCO 1kDa) and subsequently freeze-dried until further use.

Measuring degree of crosslinking using the TNBSA assay
The degree of crosslinking of CS to the primary amine groups of sECM was determined using the 2,4,6-trinitrobenzenesulfonic acid (TNBSA) assay as previously described [34,35].
Briefly, after gelation had occurred, hydrogels were incubated in 0.1M sodium bicarbonate at pH 8.5. TNBSA (0.01%) was added to the samples and incubated for two hours at 37°C under constant agitation. Sodium dodecyl sulphate (SDS, 10%) and 1M hydrochloric acid (HCl) were used to terminate and stabilize the reaction. Samples were then incubated at 120°C for 15 minutes and the absorbance of each sample was read at a wavelength of 335 nm. The concentration of free amine groups was quantified by interpolation using a standard curve of known concentrations of glycine.

Hydrogel characterisation-Swelling ratio, degradation, elution kinetics and SEM imaging
Once gelation had occurred, initial weights of hydrogels were recorded, and samples were incubated in PBS at 37°C. At different time points (1, 3, 5 and 8 hours) samples were removed from PBS, their weight was recorded and incubated again at the same conditions. The swelling ratio was calculated as the ratio of the weight of the swollen sample at different time point to that of the initial hydrogel.
The resistance of different hydrogel compositions to enzymatic digestion was evaluated using a collagenase assay [34]. Briefly, hydrogels were incubated for 1 hour in 0.1 M Tris-HCl pH 7.4 containing 50 mM CaCl2 at 37°C. Subsequently, 100 U/ml of collagenase type II (Gibco, Invitrogen, Dublin, Ireland), was added and samples were incubated at 37°C. Degradation buffer was sampled at different time points (1, 3 and 5 hrs) and stored at -20°C until biochemical analysis. Collagen content within the degradation buffer was quantified through the measurement of hydroxyproline content. Hydrogel degradation was expressed as a percentage of initial collagen detected in the sample.
Elution kinetics of CS from hydrogels was assessed under static and dynamic rotation conditions at 37°C. Briefly, after gelation, samples were transferred to 1 ml eppendorf tubes and 1 ml of PBS was added to each tube. In dynamic conditions, samples were maintained in agitation at 37°C using a rotator at 5 rpm (Stuart SB3, Cole-Parmer, UK). Native NP tissue biopsies served as a comparative control. Supernatants were assessed after 1, 3 and 6 hours and the sGAG content was determined using the diethylene blue dye binding assay. sGAG retention was calculated as a percentage of the initial sGAG content.
For SEM imaging, dehydrated samples were prepared by sputter coating with gold/palladium for 60s at a current of 40mA.

Rheological measurements and mechanical testing
Rheological measurements were performed with a MCR 102 rheometer (Anton-Paar, Hertford Herts, UK). In order to determine gelation kinetics, pre-gels with different sECM/fCS compositions were tested immediately after mixing. Storage moduli (G') and loss moduli (G'') were monitored as a function of time at a 10 Hz frequency and 1% strain at 37 o C. The point of gelation of each pre-gel composition was identified as the crossing point between storage and loss moduli curves. After gelation had occurred, a frequency sweep analysis (0-100 Hz) was performed maintaining 1% strain at 37 o C. Subsequently, the stress-realzation behavior of different gel compositions was studied by applying 10% constant strain and monitoring the relaxation over time at 37 o C.
Mechanical testing was performed on hydrogels (∅ 6 mm x h 3 mm, n=3) in unconfined compression using a single column mechanical testing machine (Zwick/Roell Z2.5, Herefordshire, UK) and a 5 N load cell. Hydrogels were maintained in a PBS bath at room temperature. A preload of 0.01 N was applied for 30s to ensure that the surfaces of the hydrogel were in contact with the loading platens and subjected to 20% strain. Stress and strain data were recorded to calculate the Young's modulus.
The hydrogel mesh size ξ (m) was determined according to the Flory equation (1): With kB representing the Boltzmann constant (J/K) and T the temperature (K) [36,37]. The storage modulus G' is bound to the Young's modulus E by the relation (2): where ν is Poisson's ratio and it is assumed to be the ideal value of 0.5, a reasonable approximation for a crosslinked gel [38]. Once combined, equations (1) and (2) define ξ as (3): 3

Cell isolation and monolayer expansion
Porcine nasal tissue from five donors of up to six months of age was obtained from a local abattoir within 24 hours of sacrifice. Cells were isolated adopting a protocol previously developed in our lab [39]. Each experiment has been performed using cells isolated from three of the five donors. Septal cartilage was finely minced to ~2mm pieces and digested with 3000 10% FBS and 2% penicillin (100 U/mL)-streptomycin (100 μg/ml) in a humidified atmosphere at 37°C and 5% O2.

Cell encapsulation, and culture
A suspension of monolayer expanded NCs was added to the pre-gel to obtain a final cell density of 1x10 6 cells/ml. 80 μl of cell seeded pre-gel were cast using a 20g gauge needle syringe into cylindrical moulds of 6mm diameter and height of 3 mm, the moulds were submerged in LG-DMEM supplemented with 10% FBS and 2% penicillin (100 U/mL)-streptomycin (100 μg/ml) and incubated for 1 hour in a humidified atmosphere at 37°C and 5% O2. After gelation, the hydrogels were incubated for 14 days in a Chemically Defined Medium (CDM) consisting of LG-DMEM supplemented with 2% penicillin (100 U/mL)-streptomycin (100 μg/ml) , 0.25 μg/ml amphotericin B, 40 μg/ml L-proline, 1.5 mg/ml bovine serum albumin (BSA), 4.7 μg/ml linoleic acid, 1% insulin-transferrin-selenium, 50 μg/ml L-ascorbic acid-2-phosphate and 100 nM Dexamethasone (all from Sigma Aldrich). Media changes were performed twice weekly and conditioned media from changes was stored at 4 o C for biochemical assays. Hydrogels were assessed through biochemical and histological analyses at days 0 and 14. Gel contraction was assessed by measuring the change in diameter of hydrogels and is expressed as a percentage decrease at each time point relative to day 0.

Determination of DNA, sulphated glycosaminoglycan and collagen content
Samples were removed from fully supplemented CDM and stored at -80°C until further analysis. Samples were subsequently digested in 500 μl each of 100 mM sodium phosphate/5mM Na2EDTA buffer containing 3.88 U/ml of papain enzyme and 5mM Lcysteine, pH 6.5 (all from Sigma-Aldrich) at 60°C under constant rotation for 12 hours. DNA content was quantified using the Hoechst Bisbenzimide 33258 dye assay (DNA QF Kit, Sigma-Aldrich). Sulfated glycosaminoglycan (sGAG) content was determined using the dimethylene blue dye binding assay. Total collagen content was quantified through the measurement of hydroxyproline content. Samples were hydrolysed at 110 o C in 38% HCl for 18 hours and assayed using a chloramine-T assay [40] and a hydroxyproline:collagen ratio of 1:7.69 [41].
Media obtained from media changes was also analysed for sGAG and collagen content.

Assessment of cell viability and morphology
Cell viability was assessed using a LIVE/DEAD® viability and cytotoxicity assay kit Semi-quantitative analysis of cell viability and cell morphology were determined using IMAGEJ software (ImageJ, National Institute of Health, Bethesda, Maryland). In order to calculate a cell's circularity coefficient, the green channel was isolated from Live/Dead images and a mask was created to replace each cell with a "best fit" ellipse. Afterwards, circularity was calculated as the ratio between the shortest and the longest axis.

Statistical analysis
Statistical analysis was performed using GraphPad Prism (version 7) software. One-way or Two-way ANOVA were used for analysis of variance. Results are displayed as mean ± standard deviation, where N represents the number of biological donors or experiments performed and n represents the technical replicates for each respectively. Significance was accepted at a level of p<0.5.

Decellularisation of ECM extracted from the NP
After completion of the treatment process from NP to ECM powder (Fig 1A), the material was decellularised and the effectiveness of the decellularisation process was assessed. Biochemical results demonstrated that the decellularisation procedure adopted successfully yielded a material whose DNA amount is below the decellularisation threshold of 50ng/mg [42] ( Fig   1B). Moreover, the length of DNA base pairs remaining in the material post-decellularisation was less than 100 bp (Fig 1C). However, while efficiently removing DNA material and maintaining the amount of collagen present in native tissue (Fig 1D, E), the decellularisation process resulted in a 96% loss of sGAG when compared to native tissue (Fig 1E, F).

Incorporation of functionalised CS in gels improves their functional behaviour
Carboxyl activated or functionalised CS interacts with amine groups on collagen fibres to form a self-crosslinkable hydrogel system (Fig 2A). The effectiveness of crosslinking was assessed through the quantification of amine groups available on collagen fibres in sECM. The concentration of free amine groups available for crosslinking significantly decreases with the addition of fCS in a concentration dependent manner (Fig 2B). Higher amounts of fCS added resulted in decreased free amine groups. Swelling capacity was also found to increase with higher amounts of fCS (Fig 2C), with increased resistance to enzymatic degradation (Fig 2D).
Gel contraction due to cellular activity was also found to decrease with fCS incorporation (82% in absence of CS; 47% for fCS containing gels), although no concentration dependence was observed ( Fig 2E). fCS containing gels also exhibited different behaviors in response to fast ramp stress-relaxation tests (Fig 2F). The group containing the highest amount of fCS achieved a higher peak stress compared to the other groups, relaxed very rapidly at short times followed by a more gradual relaxation profile. Conversely, 1:0.5 and 1:0 groups exhibited a slower relaxation curve at all times. The results from the dynamic frequency sweep revealed that G' was always greater than G", indicating that each material adopted an elastic behavior in the range of the frequencies tested, regardless of the fCS content (Fig 2G). At higher frequencies the group containing the highest amount of fCS, yielded values of storage moduli that were significantly different than the other groups and comparable to values previously reported in the literature for healthy NP tissue under similar testing conditions [43]. Moreover, the presence of higher concentrations of fCS in the gel composition resulted in shorter gelation times (Fig 2H, I).

Synthesis of sGAG and collagen are dependent on the presence of CS in ECM hydrogels
Based on preliminary testing (data not shown), the highest amount of fCS that is possible to add to the material without compromising the structural integrity of the gels was found to be at a ratio of 1:1 (2% sECM-2% fCS). We next sought to determine the influence of the fCS-ECM hydrogel on NCs using this maximal concentration (1:1) compared to a lower fCS concentration (1:0.5), to ascertain if a lower amount is sufficient to promote enhanced cell activity, compared to a control without fCS (1:0). Gels were seeded with 1x10 6 cell/ml and cultured in low oxygen (5% O2) conditions in LG-DMEM for 14 days. A significant decrease in DNA content was measured at day 14 compared to day 0 in all groups examined, although no significant difference was observed among the different groups at the same time point ( Fig   3A). Synthesis of sGAG was found to be dependent on the initial gel composition, with significantly more sGAG being deposited by cells in gels containing higher initial sGAG ( Figure 3B, C). sGAG content at day 14 was calculated by subtracting sGAG content measured at day 0. Interestingly, collagen accumulation did not appear to be influenced by differences in the initial gel composition ( Figure 3D, E). Finally, the sGAG/collagen ratio was found to increase linearly with increasing CS present in the initial gel composition (Fig 3F).

The presence of CS in the gel composition positively influences cell morphology
Motivated by previous results demonstrating a higher sGAG/collagen ratio for the gel composition of sECM:fCS of 1:1, it was decided to examine in more detail the effects of CS incorporation in sECM gels at this ratio. In order to determine the effect of CS functionalisation on cell behaviour, CS was incorporated into gels either in a functionalised (NHS+) or nonfunctionalised (NHS-) form and compared to sECM only controls. The presence of CS, both functionalised or non-functionalised, was found to positively influence the morphology of NCs inhibiting spreading and favouring the adoption of a rounded cell shape ( Figure 4A-C). Semiquantitative analysis of cell circularity corroborated these observations, demonstrating a correlation between the presence of CS and a rounded cell morphology, with a significant difference between sECM control and NHS+ group after 14 days of culture ( Fig 4D).
Interestingly, a decrease in DNA content was observed at day 14, although DNA content in NHS+ was found to be higher on day 14 compared to the control group (Fig 4F).

CS incorporation is essential to promote sGAG and collagen matrix deposition
When compared to cells in the control group, NCs cultured in CS containing gels produced significantly higher amounts of sGAG (Fig 5A), the totality of which, however, was detected in the culture media. Similar results were observed on a per cell basis when normalising sGAG per DNA content ( Fig 5B). Interestingly, when normalised by DNA content, NHS-gels yielded higher sGAG synthesis compared to NHS+ functionalised hydrogels. Histological staining demonstrated intense deposition of sGAG in the pericellular region of NCs cultured in CS containing gels at day 14, while no appreciable sGAG deposition was visible in the control group at the same time point (Fig 5C).
In terms of collagen content, NCs cultured in CS supplemented gels exhibited a reduced content at day 14 in comparison with NCs in control gels (no CS) (Fig 6A), with similar trends observed when normalised on a per cell basis ( Fig 6B). Finally, both the CS containing gels showed a significantly higher sGAG/collagen ratio than controls, the highest being the NHS-gels ( Fig   6C). Interestingly, immunohistochemical analysis revealed intense collagen type 2 and weak collagen type 1 staining in CS containing gels, while an opposite pattern was found for gels in the control group with more collagen type I being deposited compared to collagen type II ( Fig   6D). Immunohistochemical staining for collagen type 2 and collagen type 1. Scale bar is 100μm.

Functionalisation of CS improves mechanical integrity and sGAG retention of hydrogels
Mechanical integrity and stability of gels was assessed optically by comparing the degree of shape maintenance upon gelation after 10 minutes of free standing. As showed in Fig 7A, NHSgels were unable to maintain structural stability, in contrast with control and NHS+ gels.
Moreover, the elastic modulus of the three groups of gels was determined, with NHS+ gels exhibiting a significantly higher modulus compared to both NHS-and control gels (Fig 7B).
Furthermore, the elastic modulus of NHS+ gels was found to be comparable to the values previously reported in literature for native bovine NP [44,45], but still lower than native human NP [46][47][48]. To further characterise the hydrogels based on their composition, the average mesh size (ξ), corresponding to the average distance between crosslinks in the hydrogel network, was determined using equation (3). The analysis revealed a mesh size of 267.4±26.9 nm in NHS+ gels, 842.9±310.2 nm in NHS-gels, and 522.5±53.7 nm in the control group. sGAG retention was determined under dynamic and static conditions and compared to native NP. After 6 hours of dynamic agitation CS present in NHS-gels had been completely depleted, while NHS+, had retained 50% of the initial GAG content similar to native NP tissue (Fig 7C). Under static conditions, NHS+ gels retained higher levels of sGAG (>40%) when compared to NHS-gels ( Fig 7D).

Discussion
The goal of this study was to develop an injectable and self-assembled biomimetic tissue derived material and to evaluate its potential as a biomaterial for promoting nucleus pulposus regeneration. Although the use of decellularised tissue specific ECM has been widely investigated for a variety of tissue engineering (TE) applications [14,15,17,[49][50][51][52], decellularisation and solubilisation techniques inevitably result in significant loss of the sGAG [15,17,53,54]. In an attempt to reincorporate sGAG to enhance bioactivity and overcome the challenges with poor gelation in a carbohydrate rich environment, we successfully functionalised CS as a crosslinker to restore the material's biomimetic and structural characteristics. In the functionalisation process we allowed EDC to react with the carboxyl to support survival and proliferation of encapsulated articular chondrocytes [27]. Through the addition of fCS to the gel composition, we enhanced desirable characteristics of the hydrogels such as high swelling capacity and short gelation time, allowing the functionalised biomaterial to be able to hold 4 times as much water and gel twice as fast compared to sECM hydrogels.
One approach for the treatment of DD that has been commercially available for the past ten years, comprises of the injection of in vitro expanded autologous disc cells into the degenerated disc [57]. Although the study showed promising results, significant concerns have been raised in regards of cell leakage via the needle tract post-delivery [58,59]. The use of a fast-gelling material such as the one presented in this work could allow cells to be immobilised in situ in a shorter time frame and avoid leakage while simultaneously ensuring better hydration than an ECM only material.
Ideally, CS should be added to the gel composition in the same proportion as the native NP [60]. However, gel formulations with fCS:sECM ratios higher than 1 systematically failed to maintain structural stability (data not shown). The effect of CS supplementation on metabolism and differentiation has already been observed for mesenchymal stem cells (MSCs) [20,61,62] and articular chondrocytes (AC) [19,[63][64][65][66][67] for cartilage regeneration applications, however, there is little knowledge about its effect on NCs. In the first part of this study we evaluated how different gel compositions influence NC behaviour with respect to matrix deposition and found that supplementation with CS elevated the levels of matrix sGAG production. This result is consistent with previous studies reporting that CS supplementation can improve matrix accumulation for cells in either 2D or 3D culture [19,62,63,[65][66][67]. In line with our expectations, we found that the increase in sGAG production is dependent on gel composition, with a higher ratio of sGAG/Collagen being synthetized by NCs in gels prepared with higher amounts of fCS. However, although the presence of CS in the gel composition enhanced sGAG production with a sGAG/Collagen ratio comparable to the native bovine NP (typically ranging from 1:1 to 2:1), the ratio was still lower than the native human NP, which is typically 3.5:1 [3,60].
Chondrocytes are well-known to be phenotypically unstable and undergo de-differentiation when expanded in 2D monolayer or cultured in 3D collagen scaffolds or hydrogels [68][69][70].
The transition from chondrocytic to a fibroblastic phenotype has been shown to occur with a dramatic change in cell shape and metabolism, and with a marked production of fibrocartilaginous tissue, suggesting a possible relation between the two. The correlation between cell morphology and differentiation state has been under intense investigation in the past decades [71][72][73]. The presence of CS (functionalised and non-functionalised) was observed to promote a rounded cell morphology, while cells exhibited a more spread-like morphology for ECM only gels. The result suggests that the presence of CS in the cell microenvironment plays an active role in modelling its spatial conformation. There are a number of ways in which a matrix macromolecule such as CS can physiologically affect a cell: it can act as a physical barrier between the cell and the attachment sites available on the collagen structure, or act as a chemical barrier regulating ion exchanges to and from the cell [74]. It has recently been shown that GAGs can bind and sequester a wide range of proteins, the most interesting of which are the families of growth factors (GF) and cytokines. By binding these small signalling molecules, GAGs indirectly regulate cell activity by either acting as co-factor and become readily available to the cell receptor, or by limiting their availability to the cell [75,76]. The presence of CS in the cell microenvironment could therefore serve as a reservoir for cytokines and GFs to aid cell-cell communications, which could possibly help maintain their differentiated state.
In contrast with results from previous studies [66,67], the presence of CS or fCS did not have a major impact on cell viability. However, conflicting results from similar studies showed no change in cell proliferation and viability when exposed to CS, suggesting that there is still no consensus regarding the impact of sGAG on cell survival [19,63,65]. In our study we observed a decrease in the DNA levels at the end of the culture in every group, suggesting significant cell death. However, the high viability observed from Live/Dead staining at the same time point suggests that cells could have simply leaked out of the gels thereby reducing overall DNA content. This could be due to the characteristic typical of type II collagen, the main component of sECM, of forming fibre networks perhaps too loose to guarantee high cell retention during in vitro culture.
Interestingly, we found that CS supplementation, regardless of NHS functionalisation, was capable of enhancing sGAG production compared to ECM only gels while at the same time inhibiting collagen deposition in the absence of growth factor stimulation. Immunohistological analysis revealed a substantial difference in the type of collagen deposited in the pericellular region, with collagen type II being predominantly deposited in CS supplemented gels and collagen type I in ECM only gels. Differentiated chondrocytes typically produce two structural macromolecules, cartilage-specific proteoglycans and collagen type II. The transition from the synthesis of one type of collagen to another is widely known to be highly indicative of the chondrocytes' differentiation state [77][78][79]. Taken together, the results suggest that the addition of CS to a native sECM gel facilitates NCs in culture to maintain a differentiated phenotype, resulting in the deposition of the appropriate and specific ECM.
The correlation between the presence of GAGs in solution and the ability of collagen to form fibrillar reticulates has been investigated extensively over the years [23,25,80,81] demonstrating that collagen gel structures are considerably weaker when formed in the presence of GAGs. Stuart and Panitch showed in particular how fibrillogenesis in the presence of CS in a physiological environment resulted in an increase in void spaces for a collagen type I network, which yielded a reduced stiffness [81]. Interestingly, further investigations on the mechanical properties of our gel compositions revealed that when CS is supplemented without any functionalisation (NHS-) the resulting gel lacks structural stability and integrity. Another key desired characteristic is the ability to retain the incorporated sGAG. Retention kinetics demonstrated that fCS formulations were better suited to this aim and elution kinetics and mechanical properties were found to be comparable to native tissue, and superior to nonfunctionalised CS hydrogels. Overall, these findings suggest that the addition of CS in the sECM base hydrogel promotes the retention of a differentiated phenotype for NCs with concomitant specific ECM deposition, and that CS functionalisation is necessary to ensure adequate structural integrity. Preliminary in vitro culture experiments performed in our laboratory using NP cells and the fCS ECM biomaterial (data not shown), did not reveal any deleterious or negative effects in terms of cell viability or in matrix production. This is important as any injected material will come in contact with resident NP cells.

Conclusion
In this work we have successfully developed an injectable self-assembled biomimetic biomaterial combining decellularised and solubilised ECM from bovine NP tissue and functionalised CS. The inclusion of CS in the composition of the material appears to be critical for promoting a rounded cell morphology and subsequent deposition of NP-like matrix.
However, some limitations should be noted. In this study we have focused on the synthesis of sGAG as an indicator of NP-like matrix production. However, it would be important to also determine whether synthesised sGAG are integrated within a proteoglycan to reduce the likelihood that they could easily diffuse out of the organ after being synthesised. Moreover, animal studies are an important aspect for clinical translation and warrant further investigation.
Taken together, the results of this study show the potential benefit of using the proposed biomaterial in conjunction with NCs to support the regeneration of NP tissue

Funding
This work was supported by Science Foundation Ireland Career Development Award (15/CDA/3476).