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Using embedded alginate microparticles to tune the properties of in situ forming poly(N‐isopropylacrylamide)‐graft‐chondroitin sulfate bioadhesive hydrogels for replacement and repair of the nucleus pulposus of the intervertebral disc

Abstract Low back pain is a major public health issue associated with degeneration of the intervertebral disc (IVD). The early stages of degeneration are characterized by the dehydration of the central, gelatinous portion of the IVD, the nucleus pulposus (NP). One possible treatment approach is to replace the NP in the early stages of IVD degeneration with a hydrogel that restores healthy biomechanics while supporting tissue regeneration. The present study evaluates a novel thermosensitive hydrogel based on poly(N‐isopropylacrylamide‐graft‐chondroitin sulfate) (PNIPAAM‐g‐CS) for NP replacement. The hypothesis was tested that the addition of freeze‐dried, calcium crosslinked alginate microparticles (MPs) to aqueous solutions of PNIPAAm‐g‐CS would enable tuning of the rheological properties of the injectable solution, as well as the bioadhesive and mechanical properties of the thermally precipitated composite gel. Further, we hypothesized that the composite would support encapsulated cell viability and differentiation. Structure‐material property relationships were evaluated by varying MP concentration and diameter. The addition of high concentrations (50 mg/mL) of small MPs (20 ± 6 μm) resulted in the greatest improvement in injectability, compressive mechanical properties, and bioadhesive strength of PNIPAAm‐g‐CS. This combination of PNIPAAM‐g‐CS and alginate MPs supported the survival, proliferation, and differentiation of adipose derived mesenchymal stem cells toward an NP‐like phenotype in the presence of soluble GDF‐6. When implanted ex vivo into the intradiscal cavity of degenerated porcine IVDs, the formulation restored the compressive and neutral zone stiffnesses to intact values and resisted expulsion under lateral bending. Overall, results indicate the potential of the hydrogel composite to serve as a scaffold for supporting NP regeneration. This work uniquely demonstrates that encapsulation of re‐hydrating polysaccharide‐based MPs may be an effective method for improving key functional properties of in situ forming hydrogels for orthopedic tissue engineering applications.

MPs may be an effective method for improving key functional properties of in situ forming hydrogels for orthopedic tissue engineering applications. in patients is often associated with degeneration of the intervertebral disc (IVD). [3][4][5][6] The IVD is the load bearing joint between vertebrae consisting of a central nucleus pulposus (NP) and peripheral annulus fibrosus (AF). The NP is an amorphous gel comprised of collagen II and elastin fibers dispersed in a water-rich aggrecan phase. 7 In contrast, the AF has an organized, anisotropic structure made up of lamellae, or multilayered, oriented collagen fibers in an angle-ply structure. 8 The highly swellable NP expands radially under compressive loads and transfers the loads to the outer AF in circumferential tension. 9 With aging, increased catabolism reduces the collagen II and aggrecan content of the NP, 10 resulting in loss of its swelling capacity, a change in the load distribution, and the formation of tears and fissures in the AF. These structural changes may be accompanied by vascularization and neoinnervation, 11 associated with pain. 12,13 Current clinical treatments for LBP include a combination of analgesics with physical therapy or discectomy to remove nerve impinging disc tissue. 14 While these interventions provide immediate pain relief, they do not restore the healthy biomechanics to the tissue, so degeneration can continue or even accelerate. [15][16][17] Because early stage IVD degeneration is characterized primarily by changes in the NP region, the tissue is a target for newly developed therapeutic interventions. For instance, if the annulus and endplates are still competent, a swellable biomaterial can be implanted to replace the dehydrating NP. Injectability is considered paramount so the hydrogel can be implanted intradiscally with minimal damage to the AF and fill irregularly shaped tissue defects in the NP. Due to high in vivo intradiscal pressures, 18 the injectable hydrogel solution must have sufficient viscosity to be injected into an NP region without extravasation. 19 Once cured, the hydrogel should deform mechanically like the native healthy NP. 20,21 For tissue engineering approaches to repairing the IVD, the injectable hydrogel must meet these requirements and also support viability of encapsulated cells and prevent their leakage during motion and loading. 22,23 Toward these design goals, multiple in situ forming hydrogel materials have been studied for NP replacement and regeneration, like decellularized matrix-based systems, 24,25 alginate, 26 collagen, 27,28 hyaluronic acid, 29,30 and chitosan. 31 Notably, in an ex vivo test with ovine IVDs, 20 neither the implantation of hydrogel NP replacements or the re-implantation of the natural nucleus tissue restored functionality of an intact disc. It was concluded that integration with the surrounding AF tissue is a critical component of an NP replacement strategy. Thus, in parallel with the development of injectable cell-friendly systems with tunable mechanical properties, recent research has focused on engineering hydrogel bioadhesives that adhere to surrounding tissue in the IVD to minimize the risk of herniation and improve biomechanical performance. Fibrin is a biocompatible, in situ-forming carrier that forms an adhesive bond with tissue. 32 However, fibrin degrades quickly, thus making it nonideal for the long repair process of the IVD. 33 The low mechanical properties 34,35 and bioadhesive strength [36][37][38] make it inappropriate for load bearing applications. Genipin-crosslinked fibrin has been investigated as a bioadhesive cell carrier for AF repair. The covalent crosslinking of fibrin with genipin improves mechanical stiffness and adhesive strength of fibrin, 39 but genipin can have potentially cytotoxic effects. [40][41][42] An inherent challenge with fibrin-genipin is to balance the composition to improve the material properties of fibrin while promoting cell survival and extracellular matrix (ECM) deposition. 43 Injectable hydrogels based on tyramine-modified hyaluronic acid hydrogels crosslinked with horseradish peroxidase and hydrogen peroxide (H 2 O 2 ) were investigated as a bioadhesive cell carrier. However, the bonding strength to cartilage was not statistically significantly different than fibrin glue. 44 In response to these needs in NP repair, we sought to design a novel in situ forming cell carrier for NP replacement with bioadhesive properties. Previously, we reported on a novel thermosensitive graft copolymer, poly(N-isopropylacrylamide)-graft-chondroitin sulfate (PNIPAAm-g-CS). 45,46 Aqueous solutions of PNIPAAm-g-CS behave as a hydrophilic, flowable liquid below the lower critical solution temperature (LCST) of~30 C, and a precipitated, soft hydrogel above the LCST. Due to this phase transition, the copolymer can be injected into the intradiscal cavity through a small gauge needle and form a spacefilling gel in situ which is compatible with encapsulated cells. 45 The limitations of using PNIPAAm-g-CS for NP replacement are its low solution viscosity below the LCST and limited bioadhesive properties, which allow immediate extravasation upon injection into an intradiscal cavity. Therefore, we sought to improve these properties of PNIPAAm-g-CS for NP replacement by generating a composite with calcium crosslinked alginate microparticles (MPs).
Multiple levels of rationale were used for tuning PNIPAAm-g-CS properties with MPs. There is an established link between particlescale motion and rheological and mechanical properties of a suspension. [47][48][49] The viscosity of a solution increases with the addition of fine particles due to increased packing density and molecular interactions during deformation. 50 Viscosity is an important parameter for mediating adherence to tissue, since polymeric solutions that are too liquid-like lack the cohesion necessary to form substantial interactions with the tissue. 51 Embedding MPs in a hydrogel increases surface roughness, 52 which can promote mechanical fixation with a tissue substrate. Further, MPs incorporated into bulk hydrogel structures enhance network toughness. 53,54 DeVolder et al showed that poly(lactic-co-glycolic acid) MP incorporation into 3D crosslinked collagen networks increased stiffness and elasticity. 55 Notably, the encapsulation of MPs within hydrogel networks has been reported to improve mechanical performance while preserving encapsulated cell viability. 56,57 In the present study, calcium crosslinked alginate was selected to comprise the MPs because the material is inexpensive, biocompatible, hydrophilic, and can be fabricated without the use of toxic reagents. Alginate has an abundance of hydrophilic hydroxyl and carboxylic acid groups, as well as a net anionic charge, 58 facilitating swellability and physical interaction with proteins in the ECM. 59,60 Herein, we test the hypothesis that the addition of calcium crosslinked alginate MPs to aqueous solutions of PNIPAAm-g-CS would enable tuning of the rheological properties of the injectable solution, as well as the bioadhesive and mechanical properties of the precipitated gel composite, improving the suitability of the material for NP replacement. Further, we hypothesized that the composite would support encapsulated cell viability, NP differentiation, and ECM synthesis.
This study was comprised of four aims: (a) To study structureproperty relationships in injectable PNPAAm-g-CS + MP composites by varying MP concentration and diameter. Subsequently, we aimed to evaluate the formulation most closely mimicking the native NP for its ability to (b) support NP regeneration in vitro by encapsulated adipose derived mesenchymal stem cells (ADMSCs), (c) restore the degenerated porcine IVD compressive biomechanical properties ex vivo, and (d) resist expulsion from the porcine IVD cavity under lateral bending.

| MP synthesis
A water-in-oil emulsion technique was used to synthesize alginate MPs of varying diameters as described in previous publications. 45,46 Briefly, 2% (wt/vol) alginate solution (Sigma-Aldrich) and 1% (vol/vol) Tween 20 surfactant (Sigma-Aldrich) were emulsified in a vegetable oil phase. Low and high stir speeds were used to alter alginate and oil droplet size, resulting in large and small MP diameters, respectively. A 2% (wt/vol) CaCl 2 solution was added dropwise to the emulsion to crosslink the alginate. Residual oil was removed from crosslinked MPs through a series of alternating centrifugation (500g for 5 minutes) and washing steps using 70% (vol/vol) isopropanol and deionized water.
An average size was calculated for each batch by measuring the diameters of 50 randomly selected MPs. Alginate MPs were freeze dried and stored at 4 C until further use.   (Nanoscience Instruments) equipped with a cryostage. Immediately prior to analysis, the gel samples were removed from PBS, directly placed on pre-warmed foil wraps, flash frozen in liquid nitrogen, and imaged at À20 C.

| Enzymatic degradation
Approximately 0.3 mL of samples P-0 or S-50 (n = 5 per group) was immersed in 2 mL of 0.01 M PBS containing either 0.1 mg/mL collagenase P, 0.01 mg/mL hyaluronidase, 50 ng/mL aggrecanase, or 0.1 U/mL chondroitinase ABC (Sigma Aldrich). Enzyme solution was maintained at 37 C and refreshed each day for 7 days. As a control, formulations were exposed to 0.01 M PBS without enzyme. The percent mass loss was calculated using Equation (1): where, M 0 and M F are the initial and final dry masses of the sample, respectively.

| Rheological characterization
The rheological properties of each formulation were characterized Due to its bioadhesive properties, ease of injectability, and mechanical performance approaching the native NP, formulation S-50 was the focus of subsequent in vitro cell culture studies and ex vivo biomechanical tests. As a control for the cell viability analysis, metabolic activity assay, and histological characterization, cell encapsulation within P-0 was evaluated in parallel.  ADMSC metabolic activity was tracked over 14 days using the alamarBlue Cell Viability Assay (Bio-Rad). Media was removed from samples of S-50 or P-0 (n = 5 each), replaced with 300 μL of 10% ala-marBlue reagent in NP differentiation medium, and incubated for 5 hours at 37 C and 5% CO 2 . Wells without cells were used to correct for background interference. Reduced reagents were removed from the samples and absorbance readings were measured using a spectrophotometer at 570 and 600 nm. Percent reagent reduction was calculated as described by the manufacturer's instructions.

| Histology
GAG and collagen production were visualized histologically after 14 days of culture. Formulation P-0 or S-50 was fixed for 10 minutes  Antibody information is summarized in bovine, 19,64 and caprine IVD. 65 The motion segments were compressed to À1000 N and tensed to 100 N for 10 cycles at a rate of 0.1 Hz while maintained in a 37 C saline bath during testing. The peak compressive loads were scaled for differences in cross-sectional area between human and porcine and selected to represent physiological pressures of jogging or climbing stairs two at a time. 18 The first nine cycles were performed as preconditioning to establish a repeatable hysteresis response and the biomechanical parameters were calculated using the 10th cycle.
Each disc was subjected to the compression-tension cycles at each of the following conditions to detect changes in biomechanical parameters. First, the mechanical properties of the intact specimens were measured to obtain a baseline reference. Second, the specimens were punctured~15 to 30 from the coronal plane with an 18G needle ("Punctured" condition). Third, denucleation was performed using the needle attached to a syringe with vacuum ("Denucleated" condition

| Statistical analysis
Graphpad Prism 8 (San Diego, California) was used for all statistical analyses. Welch's t tests were used to identify statistical differences between sample groups. All values are reported as the mean ± SD.
Significance was set at the 95% confidence level (P < .05). The degradation behavior of formulations P-0 and S-50 in PBS and various enzymatic solutions is summarized in Figure 1C. No significant loss in dry mass between 0 and 7 days in PBS (P > .05) was measured. Exposure to the enzyme collagenase or aggrecanase did not significantly degrade the samples compared to the PBS control (P > .05). Compared to PBS, chondroitinase ABC caused a significant increase in mass loss of P-0 and S-50, at 7.6 ± 0.8% and 8.9 ± 0.8% 1.0%, respectively (P < .01). Also compared to PBS, hyaluronidase caused a significant increase in mass loss for P-0 and S-50, at 7.2 ± 0.9% and 13.8 ± 1.8%, respectively (P < .01). For hyaluronidase, significantly higher mass loss was measured for S-50 compared to P-0 (P < .01). No other enzymes produced a significantly different mass loss for P-0 compared to S-50 (P < .05).

| Adhesive properties
Histological images of the hydrogels applied to the porcine inner AF tissue substrates before adhesion testing is shown in Figure 3. Qualitative observation reveals that P-0 and fibrin hydrogel spread into a thin layer along the tissue surface, whereas S-50 retained its 3D shape comprised of a network alginate MPs.
Adhesive strength to inner AF tissue was quantitated for each formulation in tension and shear. The tensile strength of fibrin was not significantly different than P-0 ( Figure 4A

| In vitro cell culture study
After 14 days of encapsulation, ADMSCs showed excellent cellular viability within P-0 and S-50. The proportion of living cells in P-0 ( Figure 6A) and S-50 ( Figure 6B) was calculated to be 91.8 ± 1.7% and 93.4 ± 1.8%, respectively. Both P-0 and S-50 showed significant increases in reagent reduction at day 14 relative to day 0 (P < .0001), indicating cell proliferation ( Figure 6C). Reagent reduction was significantly higher for P-0 compared to S-50 at days 7 and 14 (P = .004, P < .001, respectively).
Histological staining indicated that ADMSCs seeded in P-0 and S- PCR analysis for cells encapsulated in S-50 ( Figure 9) indicate the significant upregulation of all tested major IVD ECM and NP-specific markers (P < .01 for all markers relative to day 0). Among the markers, ACAN showed the highest upregulation (≈250-fold change, Figure 9A) followed by type II collagen (≈50-fold change, Figure 9B).
Both type I collagen and SOX9 exhibited a relatively smaller upregulation (≈5-fold change, Figure 9B,C, respectively). KRT19, FOXF1, and PAX1 ( Figure 9D-F) were the highest upregulated NPspecific markers compared to HIF1α and CA12 ( Figure 9G,H). ; and H, CA12 were upregulated relative to day 0. Data were normalized to the expression levels of GAPDH. An asterisk (*) indicates a significant upregulation (P < .0001) relative to day 0. The hash symbol (#) indicates a significant upregulation (P < .01) relative to day 0 was confirmed by gross observation (Figure 10A,B) and histology ( Figure 10C-E).

| Ex vivo biomechanical testing
The axial biomechanical results are shown in Figure 11. Relative to intact, denucleation produced a significant decrease in NZ stiffness (P = .01, Figure 11A) and an increase in compressive stiffness, though not significant (P = .16, Figure 11B). The degeneration step, comprised of excessive mechanical fatigue, resulted in a statistically significant increase in compressive stiffness relative to intact (P = .03, Figure 11B). The NZ and compressive stiffnesses of the injected specimens were not significantly different than that of the intact state (P = .259 and P = .208, Figure 11A,B, respectively).
The ROM was not significantly altered from intact by injury (puncture, denucleation, degeneration) or hydrogel injection, but trended upwards with injury and downwards with implantation ( Figure 11C). The hydrogel remained within the disc space and expulsion through the annular defect was not observed with compressive-tensile loading.
Lateral bending tests were performed to evaluate the composite ability to resist expulsion from within the disc space through the needle tract. Specimens were bent to an average maximum angle of

| DISCUSSION
There is an important need for the development of injectable biomaterials that meet the requirements for NP replacement and repair.
PNIPAAm is a promising biomaterial due to its gelation behavior between room and body temperature, but the homopolymer exhibits a low water content and poor elastic properties. 67 In previous work, we demonstrated that the polymerization of NIPAAm monomer in the presence of methacrylated CS yielded a graft copolymer (PNIPAAm-g-CS), which retained the thermosensitivity of PNIPAAm with improved water retention and compressive modulus. 45 Despite the improvements, the copolymer still exhibited water and volume loss over time, limited bioadhesive properties 46 and low solution viscosity below the LCST, characteristics identified as major obstacles to successful intradiscal implantation and biomechanical performance. Thus, in the current study, we sought to improve these properties by combining PNIPAAm-g-CS with calcium crosslinked alginate MPs to form a hydrogel composite. Structure-property relationships were investigated by varying MP size and concentration. By elucidating these relationships, we sought to also shed light on the mechanism by which MPs influence the rheological, swelling, and mechanical properties of in situ forming PNIPAAm-g-CS hydrogels.
In order to prepare the bioadhesive composite for this study, dry alginate MPs were suspended in aqueous solutions of PNIPAAM-g-CS immediately prior to gelation. We postulate that when the dry MPs are suspended in solution, they begin to expand as they imbibe water and packed together to form a three-dimensional "jigsaw puzzle" within the PNIPAAM-g-CS network. This structure, discernable in the histological image in Figure 3, imparts resistance to deformation by providing a drag force within the polymer network, an effect that is evident in multiple experimental outcomes. For instance, in the rheological study, significant increases in G* were observed for all the formulations containing MPs. Notably, the drag force increases with  ranging from 3 to 5 kPa. 29 The same formulation exhibited an average confined compressive modulus of 893 kPa, only approaching the native NP tissue value of 1 MPa. 92 Last, with complex moduli G* between 1.4 and 3.5 kPa, S-50 fell short of mimicking the G* of the native NP in the same frequency range, 7.4 to 19.8 kPa. 66 Another important consideration is that the material behavior of PNIPAAm-g-CS + MPs is likely to change over time. Water is known to act as a plasticizer in hydrogels, 93 but the swelling kinetics of PNIPAAM-g-CS + MPs in situ will depend on osmotic pressure of the surrounding tissues. 94 Alginate dissolution will induce a loss of mechanical reinforcement and adhesion strength, but the rate at which this occurs depends on the ion concentration in the milieu surrounding the biomaterial. 95 Simultaneously, encapsulated cells will remodel the hydrogel network and secrete ECM, 96,97 also impacting hydrogel properties over time. Human or bovine IVD organ culture models 98-100 are the most appropriate tools for ascertaining long term hydrogel behavior within the context of an IVD-mimetic osmotic pressure, biochemical composition, and biomolecular microenvironment.
While such studies are out of the scope of the current work, it is exciting to note that the two phases in the PNIPAAm-g-CS + MP composite system can be modified to tune short and long-term behavior. For the MP phase, increasing alginate concentration would slow MP dissolution, 59 prolonging mechanical performance and bioadhesive interactions with the tissue. Another option to improve the long term bioadhesive stability of the system is to employ a recently reported two-part repair strategy, 101 where a chemically functionalized polymer layer would be placed between the bulk phase (in this case, PNIPAAm-g-CS) and surrounding AF, covalently linking the bulk phase to the tissue interface.
Despite the need for continued development, we posit that we have developed a useful platform for IVD tissue engineering. The concept of encapsulating re-hydrating polysaccharide-based MPs within a hydrogel structure can have important utility beyond the scope of this study, such as the controlled delivery of bioactive molecules for improving regenerative outcomes. From a broader perspective, we posit that the concept can be applied for improving the properties of in situ forming cell carriers in a variety of regenerative orthopedic applications.

| CONCLUSION
The Based on these results, we conclude that PNIPAAm-g-CS + alginate MPs has promise as an injectable system for NP replacement and regeneration and warrants further investigation.

ACKNOWLEDGMENTS
The work in this publication was supported by the National Institute

CONFLICT OF INTEREST
The authors declare no conflicts of interest.

DATA AVAILABILITY STATEMENT
The processed data required to reproduce these findings are available by contacting the authors.