Injectable hydrogels for cartilage regeneration

Abstract

A method for delivering stem cells for cartilage regeneration comprising substantially simultaneously administering a first composition comprising a chitosan-hydrogel precursor and a second composition comprising a hyaluronic acid-hydrogel precursor to a site in need of cartilage regeneration to provide an in situ stem cell-containing chitosan / hyaluronic acid hydrogel at the site.

Description

INJECTABLE HYDROGELS FOR CARTILAGE REGENERATIONCROSS-REFERENCE TO RELATED APPLICATION

[0001] This application claims the benefit of Application No. 63 / 735,297, filed December 17, 2024, expressly incorporated herein by reference in its entirety.BACKGROUND

[0002] Current methods for treating mild and moderate knee osteoarthritis (KOA), a chronic degenerative condition, relies on alleviating symptoms rather than promoting cartilage tissue regeneration. This is primarily because (a) cartilage has limited self-healing capabilities, due to its avascular nature and the limited proliferative potential of cartilage cells, and (b) the lack of effective regenerative medical solutions and established treatment guidelines. Stem cell therapy, a form of regenerative medicine that uses immune modulation to repair damaged tissues within the body, holds great potential for tissue regeneration. In particular, human adipose-derived stem cells (hADSCs) stand out as a favorable choice for cartilage tissue engineering. This is because (a) a large number of hADSCs can be readily isolated and harvested and (b) they possess stem cell characteristics such as self-renewal and multipotency. In the clinical setting, there is a pressing need for innovative therapeutic strategies using stem cells for both symptomatic alleviation and cartilage regeneration.

[0003] Typically, therapeutic cells are injected directly into the site of the disease using a syringe or catheter. However, less than 5% of the injected cells would be viable in the injection site a few days after the procedure. This diminished viability is largely attributed to the diffusion of the cell suspension and a decrease in cell health caused by the injective procedure and the unfavorable extracellular environment. The effectiveness of stem cell therapy in repairing damaged articular cartilage heavily relies on sustaining the presence of injected cells at the disease site over an extended period. One way to address this issue is to use a polymer hydrogel, which serves as both a cell delivery vehicle and an encapsulating matrix where the cells are confined and nurtured within the hydrogel’s highly porous structure. Synthetic polymers are commonly used for making hydrogels, because of their ease of production and property control. However, they often lack compatibility with biological systems and biodegradability when compared to natural polymer hydrogels. On the other hand, natural polysaccharides are regarded as more suitable materials for the preparation of hydrogels for biomedical applications. Polysaccharide -based hydrogelsprovide a conducive environment for cell encapsulation and support the growth of various cell types. Their large pore sizes enhance nutrient diffusion and improve cell survival. Additionally, their biodegradability reduces the need for surgical removal, and their widespread availability and cost-effectiveness further contribute to their overall advantages. Despite their advantages, current natural hydrogel systems often face challenges such as inconsistent gelation, low mechanical strength, and limited tissue retention under physiological conditions. While some have shown promise in cartilage regeneration, the mechanisms by which their material properties influence cell behavior and tissue outcomes are not well understood. This limits their clinical translation and optimization for specific regenerative applications. Additionally, injectable gels face challenges associated with crosslinking density, gelation time control, and precise cell delivery. Low-density crosslinks can result in excessive softness, whereas high-density crosslinks can lead to rigidity, risking cell death. While slow gelation causes cell dispersion and leakage, fast gelation hinders injection. The effectiveness of cell-based therapies relies on the consistent and accurate cell delivery within the hydrogel to the intended target site.

[0004] Furthermore, effective tissue adhesion is a goal for cell-laden hydrogels in knee cartilage regeneration, ensuring encapsulated cells remain at targeted sites for sustained therapeutic impact.

[0005] Despite the advantages of current natural hydrogel systems noted above, including inconsistent gelation, low mechanical strength, and limited tissue retention under physiological conditions, a need exists for improved injectable polysaccharide-based hydrogels that provide consistent and accurate cell delivery within the hydrogel to the intended target site and improved tissue adhesion for the cell-laden hydrogels to ensure encapsulated cells remain at targeted sites for sustained therapeutic impact. The present disclosure seeks to fulfill these needs and provides further related advantages.SUMMARY

[0006] In one aspect, the disclosure provides a method for delivering stem cells for cartilage regeneration. In certain embodiments, the method comprises substantially simultaneously administering a first composition comprising a chitosan-hydrogel precursor and a second composition comprising a hyaluronic acid-hydrogel precursor to a site in need of cartilage regeneration to provide an in situ stem cell-containing chitosan / hyaluronic acid hydrogel at the site, wherein the first composition, the second composition, or both the first and the second compositions further comprise stem cells.

[0007] In certain embodiments, the first composition comprises stem cells.

[0008] In certain embodiments, the second composition comprises stem cells.

[0009] In certain embodiments, the stem cells are human adipose-derived stem cells.

[0010] In certain embodiments, the chitosan-hydrogel precursor is a chitosan polymer comprising a tissue adhesive component.

[0011] In certain embodiments, the hyaluronic acid-hydrogel precursor is a hyaluronic acid polymer having aldehyde groups.

[0012] In the methods, the stem cell-containing chitosan / hyaluronic acid hydrogel is formed by Schiff base reaction between the amino groups of the chitosan polymer and the aldehyde groups of the hyaluronic acid polymer.

[0013] In certain embodiments, the hydrogel exhibits tissue-adhesive properties. In certain of these embodiments, the tissue-adhesive properties are tunable by adjusting the concentration of the chitosan-hydrogel precursor, the hyaluronic acid-hydrogel precursor, or both. In certain of these embodiments, the chitosan-hydrogel precursor composition and the hyaluronic acid-hydrogel precursor composition have polymer concentrations between 5 mg / mL and 30 mg / mL.

[0014] In certain embodiments, the hydrogel exhibits self-healing behavior following mechanical disruption. In these embodiments, the self-healing behavior arises from reversible crosslinking interactions between the chitosan and hyaluronic acid polymers. In certain embodiments, the hydrogel recovers at least 98% of its storage modulus within 3 minutes after mechanical disruption.

[0015] In certain embodiments, the hydrogel maintains stem cell viability following injection and in situ gelation.

[0016] In certain embodiments, the hydrogel supports chondrogenic differentiation of the stem cells.

[0017] In certain embodiments, the first and second compositions are administered by injection.

[0018] In certain embodiments, the site is a knee (e.g., human knee).

[0019] In another aspect, the disclosure provides a chitosan / hyaluronic acid hydrogel formed in situ in vivo, comprising one or more chitosan polymers covalently coupled to one or more hyaluronic acid polymers by Schiff base reactions, wherein the hydrogel comprises stem cells.

[0020] In certain embodiments, the stem cells are human adipose-derived stem cells.

[0021] In certain embodiments, the one or more chitosan polymers comprise a tissue adhesive component.

[0022] In certain embodiments, the one or more hyaluronic acid polymers are hyaluronic acid polymers having aldehyde groups.

[0023] In certain embodiments, the hydrogel exhibits tissue-adhesive properties. In these embodiments, the tissue-adhesive properties are tunable by adjusting the concentration of the one or more chitosan polymer, the one or more hyaluronic acid polymers, or both.

[0024] In certain embodiments, the hydrogel exhibits self-healing behavior following mechanical disruption. In these embodiments, the self-healing behavior arises from reversible crosslinking interactions between the chitosan and hyaluronic acid polymers. In certain of these embodiments, the hydrogel recovers at least 98% of its storage modulus within 3 minutes after mechanical disruption.

[0025] In certain embodiments, the hydrogels described herein maintain stem cell viability following injection and in situ gelation.

[0026] In certain embodiments, the hydrogels described herein support chondrogenic differentiation of the stem cells.

[0027] In certain embodiments, the hydrogels described herein comprises pores having an average pore size between 15 pm and 35 pm.

[0028] In a further aspect, the disclosure provides a kit for delivering stem cells for cartilage regeneration. In certain embodiments, the kit comprises (a) a first composition comprising a chitosan-hydrogel precursor, (b) a second composition comprising a hyaluronic acid-hydrogel precursor; and (c) stem cells, which may be present in the first composition, the second composition, or both.DESCRIPTION OF THE DRAWINGS

[0029] The foregoing aspects and many of the attendant advantages of this invention will become more readily appreciated as the same become better understood by reference to the following detailed description, when taken in conjunction with the accompanying drawings.

[0030] FIG. 1 is a schematic illustration of an injectable, stem cells-laden hydrogel at cartilage defect site for knee cartilage repair.

[0031] FIGS. 2A-2C show polymer chemical structures: synthesis route and chemical structure of a N-succinyl chitosan (Suc-CS), a representative chitosan polymer useful for making the in situ hydrogel described herein (FIG. 2A); synthesis route and chemical structure of an aldehyde hyaluronic acid (Ald-HA), a representative hyaluronic acid polymer useful for making the in situ hydrogel described herein (FIG. 2B); and the chemical structure of crosslinked Suc-CS / Ald-HA hydrogel, a representative in situ hydrogel described herein (FIG. 2C).

[0032] FIGS. 3A and 3B compare gelation capability of representative hydrogels: rheological analysis, with time-sweep mode, of various hydrogel samples including Ald- HA, Suc-CS, and Suc-CS / Ald-HA, post mixing from 0 to 1200 s (temperature: 37 °C, shear strain: 1%, and frequency: 1 Hz) (FIG. 3A) and rheological analysis, with time-sweep mode, of Suc-CS / Ald-HA hydrogels post mixing from 0 to 50 s (temperature: 37 °C, shear strain: 1%, and frequency: 1 Hz) (FIG. 3B). Each polymer (Suc-CS or Ald-HA) was dissolved in DPBS at a concentration of 20 mg / niL. Suc-CS / Ald-HA hydrogels were prepared by mixing same volume of Suc-CS and Ald-HA solutions.

[0033] FIGS. 4A-4F compare mechanical strengths of representative hydrogels measured by rheology. The storage (G') and loss modulus (G”) were recorded under increasing shear stress from 0.1-10000 Pa for Suc-CS / Ald-HA hydrogels with different concentrations (10, 20, 30, 40, and 50 mg / mL) (temperature: 37 °C, frequency: 1 Hz) (FIGS. 4A-4E, respectively). The storage modulus (G’) within linear viscoelastic regime of Suc-CS / Ald-HA hydrogels with different concentrations (10, 20, 30, 40, and 50 mg / mL) (FIG. 4F).

[0034] FIG. 5 illustrates the porous structures of representative hydrogels. SEM images of freeze-dried Suc-CS / Ald-HA hydrogels prepared at polymer concentrations of 10, 20, 30, 40, and 50 mg / mL and incubated in DI water at 37 °C for 1, 3, and 7 days. The short scale bar represents 20 pm, and the long scale bar represents 100 pm.

[0035] FIG. 6 illustrates the pore diameters of representative hydrogels. The pore diameters of the freeze-dried Suc-CS / Ald-HA hydrogels prepared at polymer concentrations of 10, 20, 30, 40, and 50 mg / mL and incubated in DI water at 37 °C for 1, 3, and 7 days. The pore diameter was obtained by evaluating SEM images using ImageJ. Two-way ANOVA was used, followed by Tukey’s multiple comparisons test (mean ± SD, n > 10; * p < 0.05, ** p < 0.01, **** p < 0.0001; ns: nonsignificant).

[0036] FIGS. 7A-7D compare the diameters of cell spheroids assessed from day 1 to day 14. The hADSCs were encapsulated in hydrogels prepared at concentrations of (a) 20 mg / mL, (b) 30 mg / mL, (c) 40 mg / mL, and (d) 50 mg / mL (FIGS. 7A-7D, respectively). One-way ANOVA was used, followed by Tukey’s multiple comparisons test (mean ± SD, n > 5; * p < 0.05, ** p < 0.01, *** p < 0.001, **** p < 0.0001; ns, nonsignificant).

[0037] FIGS. 8A-8C compare gene expression (fold changes) associated with chondrogenic differentiation. Relative expression of RNA: ACAN (FIG. 8A), SOX9 (FIG. 8B), and COL2A1 (FIG. 8C) contents in hADSCs cultured on 2D TCPS in 3D hydrogels prepared at 20 mg / mL, and in 3D hydrogels prepared at 30 mg / mL in the chondrogenic differentiation medium for a 14-days culture period. Gene expression was first normalized to glyceraldehyde 3-phosphate dehydrogenase (GAPDH) (reference gene) and then normalized to the expression associated with 2D TCPS culture on day 0 (set as 1 - time). One-way ANOVA followed by Tukey’s multiple comparison test was used to determine whether the differences were statistically significant (mean ± SD, n = 3; ** p < 0.01, *** p < 0.001, **** p < 0.0001; ns, nonsignificant).

[0038] FIG. 9 is a schematic illustration of a representative stem cells-laden hydrogel with various functions for knee cartilage repair.

[0039] FIGS. 10A and 10B illustrate the synthesis route and chemical structure of hydroxybutyl chitosan (HBCS) and catechol-functionalized HBCS (Dopa-HBCS), representative chitosan polymers useful for making the in situ hydrogels described herein (FIG. 10A), and the synthesis route and chemical structure of aldehyde hyaluronic acid (Ald-HA), a representative hyaluronic acid polymer useful for making the in situ hydrogels described herein (FIG. 10B).

[0040] FIG. 11 illustrates the chemical structure of a representative Ald- HA / Dopa-HBCS hydrogel.

[0041] FIGS. 12A-12F compare mechanical strength and gelation kinetics of representative hydrogels: rheological evaluation of hydrogel mechanical strength and gelation kinetics. Shear sweep tests measured storage modulus (G’) and loss modulus (G”) across a shear stress range of 0.1-1000 Pa for Ald-HA / Dopa-HBCS mixtures with fixed volume ratios but varying concentrations: (a) 5, (b) 10, (c) 20, and (d) 30 mg / mL (FIGS. 12A-12D, respectively). Tests were also conducted with volume ratios of 7 / 3, 5 / 5, and 3 / 7 at 10 mg / mL in DI water (37°C, 1 Hz), stopping near the linear viscoelastic regime (LVR)limit where structural breakdown begins (FIG. 12E). Time-sweep tests monitored gelation, recording G’ and G” for 5 / 5 and 3 / 7 ratios at 10 mg / mL in DI water over 0-20 minutes (37°C, 1% strain, 1 Hz) (FIG. 12F). (+) denotes DI water with 100 mM CaCh; (-) denotes pure DI water. In sample names, A is Ald-HA, D is Dopa-HBCS, and H is HBCS, with numbers indicating volume ratios at equal concentrations (e.g., A5H5: Ald-HA and HBCS; A5D5: Ald-HA and Dopa-HBCS).

[0042] FIGS. 13A-13C illustrate representative hydrogel morphology and degradation. SEM images of freeze-dried Ald-HA / Dopa-HBCS hydrogels of different volume mixing ratios in DI water containing 100 mM CaCh (FIG. 13 A). All scale bars represent 50 pm. Pore diameter of freeze-dried Ald-HA / Dopa-HBCS hydrogels of different volume mixing ratios in DI water containing 100 mM CaCh (FIG. 13B). Pore diameter was obtained through SEM image analysis using Imagel. Statistical significance was evaluated using one-way ANOVA followed by Tukey’s multiple comparisons test (mean ± SD, n > 15; **** p < 0.0001; ns = nonsignificant). Weight remaining ratio of lyophilized Ald-HA / Dopa-HBCS hydrogels of different volume mixing ratios at 10 mg / mL in complete mesenchymal stem cells culture medium containing 100 mM CaCh over time (0, 1, 3, 5, 7, 9, and 14 days) (FIG. 13C). A5D5 and A3D7 denote Ald-HA and Dopa- HBCS mixtures of volume ratios of 5 / 5 and 3 / 7, respectively. A5H5 and A3H7 denote Ald- HA and HBCS mixtures of volume ratios of 5 / 5 and 3 / 7, respectively.

[0043] FIG. 14 compares the effect of culture environment on hADSC proliferation obtained through the Alamar Blue assay. Cell proliferation of hADSCs cultured in four different hydrogels: A5D5, A3D7, A5H5 and A3H7 on day 1, 3, 5, 7, and 14. A5D5 and A3D7: Ald-HA and Dopa-HBCS mixtures at volume mixing ratios of 5 / 5 and 3 / 7 at 10 mg / mL in DI water with 100 mM CaCh. A5H5 and A3H7: Ald-HA and HBCS mixtures at volume mixing ratios of 5 / 5 and 3 / 7 at 10 mg / mL in DI water with 100 mM CaCh-

[0044] FIGS. 15A-15C present the evaluation of representative hydrogel tissue adhesion strength. Schematic representation of catechol-mediated tissue adhesion via covalent bonding mechanisms (FIG. 15A); adhesive strength profiles of three representative hydrogels (A5H5, A5D5, and A5DH5) on porcine skin plotted against applied strain (FIG. 15B); and the maximum adhesive strength of three representative hydrogels (A5H5, A5D5, and A5DH5) on porcine skin (FIG. 15C). A5D5: Ald-HA and Dopa-HBCS mixture at a volume mixing ratio of 5 / 5 at 10 mg / mL in DI water with 100mM CaCb. A5H5: Ald-HA and HBCS mixture at a volume mixing ratio of 5 / 5 at 10 mg / mL in DI water with 100 mM CaCh. A5DH5: Ald-HA, Dopa-HBCS, and HBCS mixture at a volume mixing ratio of 5 / 2.5 / 2.5 at 10 mg / mL in DI water with 100 mM CaCh. Statistical analysis was performed using one-way ANOVA with Tukey’s post hoc test to assess the significance of differences among groups (mean ± SD, n=5, *p < 0.05; **p < 0.01).

[0045] FIGS. 16A-16C compare expression of chondrogenic markers in different culture conditions. Quantitative RT-PCR analysis of gene expression in hADSCs cultured for 14 days under various conditions: ACAN (FIG. 16A), SOX9 (FIG.16B), and COL2A1 (FIG. 16C). Cells were maintained either on 2D TCPS or in 3D hydrogels, using either standard growth medium or chondrogenic differentiation medium. Gene expression levels were first normalized to the housekeeping gene GAPDH, then to the baseline expression on 2D TCPS at day 0 (set to 1). Statistical analysis was conducted using one-way ANOVA followed by Tukey’s post hoc test (mean ± SD, n = 3 *p < 0.05, **p < 0.01, ***p < 0.001, **** / ? < 0.0001 ; ns = not significant). Abbreviations: 2D - cells on TCPS in growth medium; 2D-d - cells on TCPS in differentiation medium; 3D-d - cells in 3D hydrogels in differentiation medium.DETAILED DESCRIPTION

[0046] The present disclosure provides injectable, biodegradable, self-healing, and in situ forming polysaccharide hydrogels, as a matrix that encapsulates and maintains the viability of human adipose-derived stem cells (hADSCs) for cartilage regeneration (FIG. 1). In one embodiment, the hydrogel is composed of two natural polysaccharide polymers, a chitosan bearing amino groups (Suc-CS) and aldehyde HA (Ald-HA). The carbon-carbon bonds of HA were cleaved through an oxidation reaction, yielding new aldehyde functional groups transforming it into Ald-HA. The aldehyde groups of Ald-HA then reacted with the amino groups of Suc-CS via Schiff base reaction, leading to the formation of a covalent crosslinked hydrogel network. The injectable solutions of these two natural polysaccharide polymers were prepared separately and mixed during injection using a double -barrel syringe to create the hydrogel. The hADSCs were incorporated in the Ald-HA solution and became encapsulated in the hydrogel during the gelation process. The disclosure describes the effect of polymer concentrations on the gelation capacity, the physical properties of the hydrogel (such as mechanical strength and pore size), as well as the morphology and viability of the encapsulated hADSCs over time. Furthermore, thetemporal changes in gene and protein expression levels associated with chondrogenic differentiation within the encapsulated hADSCs were examined.

[0047] The method provides for the delivery of stem cells directly to a site in need of cartilage regeneration. In the method, the substantially simultaneous delivery of a first composition comprising a chitosan-hydrogel precursor and a second composition comprising a hyaluronic acid-hydrogel precursor provides a stem cell-containing chitosan / hyaluronic acid hydrogel at the site. The substantially simultaneous delivery of the first composition and the second composition produces the stem cell-containing chitosan / hyaluronic acid hydrogel in situ. Prior to the substantially simultaneous delivery at the site, the first composition and the second composition are separate (i.e., not in contact, administered independently) and are combined at the site to produce the hydrogel via reaction of the components of the first and second compositions.

[0048] In certain embodiments, the substantially simultaneous delivery of the first composition and the second composition is achieved using a double -barrel syringe that enables simultaneous injection and rapid in situ gelation, maintaining stem cell viability and ensuring a consistent, controlled delivery directly at the site of application. This approach offers enhanced precision and practicality in clinical settings, as it allows for minimally invasive administration and avoids complications from mixing outside the body. Such a method would be challenging to replicate through conventional injection means, as they lack the capability for controlled, synchronized delivery.

[0049] In the method, two hydrogel precursor compositions are delivered to the site where they combine to form the hydrogel in situ. The first composition includes a chitosan-hydrogel precursor (i.e., a chitosan polymer suitable for hydrogel formation with a hyaluronic acid). The second composition includes a hyaluronic acid chitosan-hydrogel precursor (i.e., a hyaluronic acid polymer suitable for hydrogel formation with a chitosan polymer). In the practice of the present method, the in situ stem cell-containing chitosan / hyaluronic acid hydrogel is formed by Schiff-base crosslinking of the chitosan and hyaluronic polymers. A Schiff-base crosslink is formed between an amino group of the chitosan polymer and an aldehyde group of the hyaluronic acid. The degree of Schiff- base crosslinking affects the overall properties of the hydrogel, including its mechanical strength, as described herein. For example, the greater the extent of crosslinking (hydrogel crosslinking density), the greater the hydrogel’s mechanical strength.

[0050] The Schiff-base crosslinking advantageously imparts self-healing characteristics to the hydrogel. The Schiff-base reaction facilitates dynamic crosslinking, making the hydrogel resilient and adaptable to mechanical stresses. This feature is advantageous for applications requiring long-term durability within the body and offers a unique advantage in tissue regeneration contexts.

[0051] In addition to the chitosan and hyaluronic acid polymers, the first and / or second compositions further include stem cells suitable for cartilage regeneration in situ. The stem cells of the first and / or second compositions are encapsulated in the hydrogel on formation of the hydrogel in situ.

[0052] FIGS. 1 and 9 are schematic illustrations of representative methods for delivering stem cells for cartilage regeneration as described herein.

[0053] The chitosan-hydrogel precursor of the first composition is a chitosan polymer suitable for hydrogel formation with a hyaluronic acid.

[0054] The chitosan polymer has a molecular weight (number-average Mn or weight-average Mw, e.g., as determined by SEC / GPC or an equivalent method) in the range of about 50 kDa to about 350 kDa. A preferred molecular weight range is about 190 kDa to about 310 kDa. In some embodiments, the chitosan has a corresponding degree of polymerization (DP) of about 300 to about 2000, with a preferred DP range of about 900 to about 1500.

[0055] In certain embodiments, the chitosan polymer is a succinylated chitosan polymer. A representative succinylated chitosan polymer (e.g., Suc-CS) is illustrated in FIG. 2A.

[0056] The succinylated chitosan polymer (Suc-CS) has a degree of substitution (DS) corresponding to the succinic acid (succinyl) content. As used herein, “DS” refers to the molar percentage of chitosan glucosamine repeat units functionalized with a succinyl group. In some embodiments, the DS is in the range of about 5% to about 80%, preferably about 20% to about 50%, and more preferably about 25% to about 40%. In one non-limiting example, the Suc-CS has a DS of about 34%, as determined byNMR by comparing the integral of the CH2-CH2 protons of the succinyl group (near 2.5 ppm) with the integral of the methyl protons (around 2.0 ppm).

[0057] In certain embodiments, the chitosan polymer includes a tissue adhesive component to advantageously enhance adhesion of the in situ hydrogel to the surrounding tissues. In certain of these embodiments, the tissue adhesive component is a3,4-dihydroxyphenylacetic acid (Dopa) group. In certain embodiments, the chitosan polymer including the tissue adhesive component is prepared from a hydroxybutylated chitosan polymer (e.g., HBCS). A representative chitosan polymer that includes a tissue adhesive component (e.g., Dopa-HBSC) is illustrated in FIG. 10A.

[0058] In some embodiments, the hydroxybutyl chitosan (HBCS) comprises hydroxybutyl substituents at a degree of substitution (DS) of about 40% to about 85% (e.g., mol% of chitosan repeats units bearing hydroxybutyl groups). In preferred embodiments, the DS is about 55% to about 75%, and in more preferred embodiments, the DS is about 60% to about 70%. In one exemplary embodiment, HBCS synthesized from medium molecular weight chitosan (about 190-310 kDa) exhibited a DS of about 65.3%, as determined by ’H NMR integration comparing the anomeric protons (Hl, 4.5-4.8 ppm) to the hydroxybutyl methyl protons (-CH3, 0.9 ppm).

[0059] In some embodiments, the catechol-modified hydroxybutyl chitosan (Dopa-HBCS) comprises a catechol (Dopa) degree of substitution (DS) (i.e., mol% catechol groups relative to the chitosan repeat units) of about 1 % to about 30%. Preferably, the catechol DS is about 5% to about 20%; more preferably, about 10% to about 18%. In certain embodiments, Dopa-HBCS is prepared by carbodiimide-mediated coupling of a catechol-containing acid to HBCS, where HBCS is synthesized from medium molecular weight chitosan (190-310 kDa). In one representative embodiment, new peaks at about 6- 7 ppm in the 'H NMR spectrum confirm catechol protons, and the catechol DS is about 14.9%.

[0060] The hyaluronic acid-hydrogel precursor of the second composition is a hyaluronic acid polymer suitable for hydrogel formation with a chitosan polymer.

[0061] The hyaluronic acid polymer (e.g., sodium hyaluronate) has a molecular weight (number-average Mn or weight-average Mw, e.g., as determined by SEC / GPC, SEC-MALS, intrinsic viscosity, or an equivalent method) in the range of about 1.0 x 106Da to about 2.0 x 106Da. A preferred molecular weight range is about 1.5 x 106Da to about 1.8 x 106Da (e.g., hyaluronic acid sodium salt derived from Streptococcus equi). In some embodiments, the hyaluronic acid has a corresponding degree of polymerization (DP), defined as the number of disaccharide repeat units, of about 2400 to about 5000, with a preferred DP range of about 3600 to about 4500 (assuming an average disaccharide repeat-unit molecular weight of -401 g / mol for sodium hyaluronate).

[0062] In certain embodiments, the hyaluronic acid polymer is an aldehyde- containing hyaluronic acid polymer. A representative aldehyde-containing hyaluronic acid polymer (e.g., Ald-HA) is illustrated in FIGS. 2B and 10B. The aldehyde-containing hyaluronic acid polymer is prepared by oxidation of hyaluronic acid by, for example, treatment with periodic acid. The aldehyde content of the aldehyde-containing hyaluronic acid polymer can be adjusted by the oxidation conditions to impart the desired aldehyde content to the hyaluronic acid, which in turn influences the degree of Schiff-base crosslinking with the chitosan polymer in the product in situ hydrogel.

[0063] The aldehyde-functionalized hyaluronic acid (Ald-HA) has an aldehyde content, expressed as a degree of substitution (DS) or oxidation percentage (i.e., mol% of HA repeat units converted to aldehyde-bearing units) in the range of about 5% to about 30%. A preferred aldehyde content range is about 10% to about 20%. In one representative embodiment, the oxidation percentage is about 14%, as quantified by measuring the number of aldehyde groups in the polymer using a tert-butyl carbazate assay.

[0064] The in situ stem cell-containing chitosan / hyaluronic acid hydrogel is produced by combining the first composition and the second composition at the site in need of cartilage regeneration.

[0065] The chitosan polymer solution and the hyaluronic acid polymer solution may be combined and delivered at a volumetric mixing ratio (chitosamHA, v / v) in the range of about 1:9 to about 9:1. A preferred mixing ratio range is about 1 :3 to about 3:1, and in certain embodiments the ratio is about 1: 1. In some embodiments, the total polymer concentration (combined concentration of chitosan polymer and hyaluronic acid polymer in the delivered mixture) is about 5 mg / mL to about 60 mg / mL, preferably about 10 mg / mL to about 40 mg / mL. In representative embodiments, the total polymer concentration is about 10 mg / mL or about 20 mg / mL.

[0066] The in situ chitosan / hyaluronic acid hydrogel comprises dynamic Schiff- base (imine) crosslinks formed between aldehyde groups on aldehyde-functionalized hyaluronic acid (Ald-HA) and primary amine groups on the chitosan polymer. As used herein, “Schiff-base crosslink density” may be expressed as an equivalent measure based on the stoichiometric ratio of aldehyde groups to amine groups (Ald:NH2, on an equivalent basis) in the precursor solutions and / or the fraction of functional groups participating in imine formation. In some embodiments, the Ald:NH2 equivalent ratio is in the range of about 0.05: 1 to about 5: 1, preferably about 0.2:1 to about 2:1, and more preferably about0.5:1 to about 1.5:1. In some embodiments, the effective crosslink density corresponds to about 1% to about 30% of polymer repeat units participating in Schiff-base linkage formation, preferably about 5% to about 20%, and more preferably about 10% to about 20%, depending on the aldehyde content of Ald-HA and the available amine content of the chitosan polymer.

[0067] A representative chitosan / hyaluronic acid (Suc-CS / Ald-HA) hydrogel is illustrated in FIG. 2C. Tn FIG. 2C, 1, m, and n represent the mole percent (mol%) of repeat units in the succinylated chitosan (Suc-CS) backbone, where 1 + m + n = 100 mol%. In some embodiments, 1 (N-acetylated chitosan units) is about 1-30 mol%, preferably about 5-20 mol%. In some embodiments, m (succinyl-substituted chitosan units) is about 5-60 mol%, preferably about 20-50 mol%, and more preferably about 30-40 mol% (e.g., about 34 mol%). In some embodiments, n (unsubstituted / deacetylated chitosan units) is about 10-94 mol%, preferably about 30-70 mol%. In FIG. 2C, k represents the mole percent (mol%) of oxidized hyaluronic acid repeat units bearing aldehyde groups (i.e., aldehyde functionalization level of Ald-HA). In some embodiments, k is about 5-30 mol%, preferably about 10-20 mol%, and in certain embodiments about 14 mol%.

[0068] A representative chitosan / hyaluronic acid having a tissue adhesive component (Dopa-HBCS / Ald-HA) hydrogel is illustrated in FIG. 11. In FIG. 11, a, b, and c denote the relative amounts of repeat units in the Dopa-HBCS polymer bearing the indicated substituents, expressed as mole percent (mol%) relative to total chitosan repeat units (i.e., an equivalent measure of degree of substitution). In some embodiments, a (catechol / Dopa-substituted units) is about 1-30 mol%, preferably about 5-20 mol%, and more preferably about 10—18 mol% (e.g., about 14.9 mol%). In some embodiments, b (hydroxybutyl-substituted units) is about 40-85 mol%, preferably about 55-75 mol%, and more preferably about 60-70 mol% (e.g., about 65.3 mol%). In some embodiments, c (N- acetylated units) is about 1-30 mol%, preferably about 5-20 mol%. In FIG. 11, n represents the mole percent (mol%) of oxidized hyaluronic acid repeat units bearing aldehyde groups (i.e., aldehyde functionalization level of Ald-HA), which is about 5-30 mol%, preferably about 10-20 mol%, and in certain embodiments about 14 mol%.

[0069] Chitosan / Hyaluronic Acid-Based Hydrogels

[0070] In one aspect, the disclosure provides chitosan / hyaluronic acid-based hydrogels.

[0071] Current knee osteoarthritis (KOA) treatments mainly provide symptom relief rather than cartilage repair. While regenerative medicine using stem cell therapy holds promise for tissue regeneration and joint function restoration, a significant challenge lies in the efficient and minimally invasive delivery of stem cells to target sites and ensuring high regenerative efficacy. This challenge stems from issues such as cell leakage and reduced cellular activity post-transplantation. An injectable polysaccharide hydrogel is provided that is compatible with cells and tissues and is suitable to support the proliferation of human adipose-derived stem cells (hADSCs) for cartilage regeneration. The hydrogel is formed on-site at the defect site of articular cartilage by mixing two injectable polymer solutions at physiological temperature post-injection. During the gelation process, hADSCs contained in one of the polymer solutions are encapsulated in the hydrogel. The hydrogel is tailored to create a desired microenvironment with mechanical properties, pore size, and degradation rate suitable for supporting hADSC viability and function. Nearly all of the encapsulated hADSCs remained viable 14 days post-injection and exhibited increased expression of chondrogenic differentiation genes compared to those cultured on 2D surfaces.

[0072] The preparation, characterization, and use of a representative chitosan / hyaluronic acid-based hydrogel (Suc-CS / Ald-HA hydrogel) is described below.

[0073] Suc-CS and Ald-HA polymer synthesis and characterization

[0074] The synthesis of Suc-CS involved the introduction of succinyl groups to the chitosan’s N-tenninal glucosamine units. The medical grade chitosan, characterized by a 95% deacetylation degree, was dissolved in a 5% (v / v) acetic acid solution, followed by the addition of methanol. The succinic anhydride was dissolved in acetone and added to the above solution. The reaction was allowed to proceed under stirring at room temperature for 24 h. The polysaccharide amine group and the electrophilic carbonyl group of the anhydride underwent a condensation process during the succinylation reaction, which resulted in the formation of an amidic bond and the opening of the anhydride ring. The synthesis route and chemical structure of Suc-CS are shown in FIG. 2A. The synthesized Suc-CS exhibited excellent solubility in Dulbecco’s phosphate -buffered saline (DPBS) at room temperature at concentrations not exceeding 50 mg / mL, confirming the successful conjugation of the succinyl group. The substitution degree of Suc-CS was 34%, which was calculated by comparing the integral of the peaks of the CH2-CH2 section of thesuccinyl group on Suc-CS near 2.5 ppm and the integral of the peaks of the methyl protons around 2.0 ppm.

[0075] To synthesize Ald-HA, HA was dissolved in deionized (DI) water, and sodium periodate was then introduced to initiate the reaction at room temperature. By reacting with sodium periodate, which oxidizes the vicinal hydroxyl groups to dialdehydes and opens the sugar ring to create dialdehyde derivatives, aldehyde groups were incorporated into the hyaluronic acid. The synthetic route and chemical structure of aldehyde hyaluronic acid are shown in FIG. 2B. The presence of a peak around 5 ppm in the spectrum indicates the presence of aldehyde groups in HA. The oxidation percentage was quantified at 14% by measuring the number of aldehydes in the polymer using tertbutyl carbazate. This was determined by comparing the peak corresponding to the tertbutyl substituent ((CH^COCONHNH-. 5 = 1.38 ppm) with the peak of HA acetamide protons at 1.9 ppm. It is worth noting that the preparation process of hydrogel is facilitated by a substantial reduction in the viscosity of Ald-HA, which results from the disruption of intermolecular hydrogen bonds during the oxidation of HA.

[0076] Suc-CS / Ald-HA hydrogel characterization

[0077] The hydrogels were prepared by gently mixing the solutions of the two modified polysaccharides, Suc-CS and Ald-HA. Both Suc-CS and Ald-HA were dissolved in DPBS at concentrations of 10, 20, 30, 40, and 50 mg / mL, respectively. The gelation mechanism is thought to be based on the Schiff base reaction between the amino of Suc- CS and aldehyde groups of Ald-HA (FIG. 2C). To confirm gelation, rheological analysis was performed using the time-sweep mode immediately after mixing. To establish the appropriate mechanical strain and frequency settings for time-sweep experiments, the gelled samples were initially subjected to incremental strain measurements at a frequency of 1 Hz and a temperature of 37 °C. The result showed that strains ranging from 0.1% to 5% fall within the linear viscoelastic region (LVR), where the storage modulus remains unaffected by the applied strain. Consequently, a strain of 1% at 1 Hz was selected as the optimal testing condition to minimize both gel damage and noise during the subsequent time-sweep experiments. FIG. 3A shows the changing of the storage (G’) and loss (G”) modulus of Ald-HA (20 mg / mL, DPBS), Suc-CS (20 mg / mL, DPBS), and Suc-CS / Ald- HA mixture (20 mg / mL, DPBS) over time at 37 °C. The storage modulus (G’) represents the solid-like behavior of the sample, indicating its elastic characteristics, while the loss modulus (G”) represents the liquid-like behavior, indicating its viscous characteristics. Aviscoelastic solid hydrogel exhibits a higher storage modulus than loss modulus (G’ > G”) due to the crosslink inside the hydrogel, while a viscoelastic liquid solution exhibits a lower storage modulus than loss modulus (G’ < G”) because of the lack of strong interactions between polymer chains. FIG. 3A shows that the storage modulus was larger than loss modulus (G’ > G”) for the Suc-CS / Ald-HA mixture, confirming the hydrogel formation. On the other hand, both Ald-HA and Suc-CS solutions had a storage modulus smaller than the loss modulus (G’ < G”), indicating that these individual polymer solutions could not form hydrogels on their own. To demonstrate the occurrence of the Schiff base reaction during crosslinking of Ald-HA and Suc-CS solution, Suc-CS was mixed with HA that lacked aldehyde groups (Suc-CS / HA). The mixture was then left at room temperature overnight after mixing using a vortex to ensure complete interaction between polymer chains. Subsequently, the storage (G’) and loss (G”) modulus of the fully mixed Suc- CS / HA were measured using rheological analysis in stress sweep mode (unlike the time sweep mode, as the gelation time was not known). The results show that the storage modulus was lower than the loss modulus (G’ < G”), indicating no Schiff base reaction between Suc-CS and HA. Thus, the Suc-CS / Ald-HA hydrogel could be readily prepared by mixing two polymer solutions at the same concentrations (1 / 1, v / v) at 37 °C.

[0078] The preparation of an injectable hydrogel can be achieved through two primary strategies. The first strategy involves in situ gelation of liquid polymers postinjection, wherein precursor solutions are combined, and gelation occurs upon injection into the target site. This method enables a controlled and localized formation of the hydrogel within the body. Alternatively, the second approach employs shear-thinning polymers. In this case, the hydrogel precursor experiences a temporary reduction in viscosity under shear stress during injection, facilitating smooth administration. Subsequently, the polymer system regains its original viscosity after injection, resulting in the formation of a stable hydrogel at the desired location. As described herein, the hydrogel was created by gently mixing solutions of modified polysaccharides, specifically Ald-HA and Suc-CS utilizing the first strategy. The gelation was confirmed by the rheological analysis using the time-sweep mode immediately following injection. The intersection points of the storage (G’) and loss (G”) modulus was defined as the gelation time. The results of rheological analysis for Suc-CS / Ald-HA hydrogel in FIG. 3B showed that the Suc-CS / Ald-HA mixture exhibited solid hydrogel behavior as early as 10 s after mixing, guaranteeing immediate encapsulation of hADSCs during the injection process.

[0079] Building on the foundational property of rapid gelation, the injectability of a hydrogel also plays a role in delivering therapeutic agents to the target site with minimal invasiveness. For example, the injection of Suc-CS / Ald-HA hydrogel can be made through a medical grade double-barrel syringe (Duploject syringe, Baxter) fitted with a 21 G needle (Becton Dickinson), allowing for varying concentration from 10 to 50 mg / mL. This feature ensures that the injection of our hydrogel can be optimized according to the specific requirements of the treatment sites.

[0080] Self-healing hydrogels have found wide applications in the delivery of cells, drugs, biomolecules, and polynucleotides. These hydrogels exhibit the remarkable ability to restore their structural integrity and mechanical properties (e.g., elasticity, viscosity, stiffness, and toughness) under stress. The rapid self-healing capability of representative Suc-CS / Ald-HA hydrogels was demonstrated; mending within 3 min after two separated pieces of hydrogel were brought together. The self-healing mechanism of Suc-CS / Ald-HA hydrogels relies on the presence of dynamic covalent bonds, particularly imide bonds formed via the Schiff base reaction. Moreover, the self-healing capability of this hydrogel would allow it to transform into the shape of and fill the defect site after injection. This property can significantly enhance its therapeutic efficacy, as it ensures complete coverage of the affected area, promoting better tissue repair and regeneration.

[0081] Rheological analyses were performed on samples after the crosslinking of Suc-CS and Ald-HA to evaluate the effect of the mechanical properties of the hydrogels on stem cell proliferation and phenotype. The modulus of the Suc-CS / Ald-HA hydrogel was assessed through stress sweep testing at 37 °C at different polymer concentrations (10, 20, 30, 40, and 50 mg / mL). For all the Suc-CS / Ald-HA hydrogels prepared at different concentrations and tested at 37 °C, it was observed that the storage modulus (G’) was higher than the corresponding loss modulus (G”) (FIGS. 4A-4E). This observation indicated the successful gelation of the Suc-CS and Ald-HA solutions across all concentrations. As the polymer concentrations increased, the storage modulus (G’) within the linear viscoelastic regime (LVR) also increased, ranging from 75 Pa to 5000 Pa (FIG. 4F), due to the increased crosslinking density. The increased crosslinking enhanced the elasticity and allowed the hydrogel to maintain structural integrity even under higher stresses. Notably, the hydrogel matrix stiffness could be tuned by varying the polymer concentration. It is worth mentioning that the Suc-CS and Ald-HA solutions, prepared at different concentrations, formed a hydrogel within just 1 min after merely mixing twopolymer solutions at the same concentrations (1 / 1, v / v) at 37 °C. This rapid gelation capability would enable the in situ gelation immediately after mixing and allow for the encapsulation of hADSCs contained in the Ald-HA solution in the hydrogel upon injection.

[0082] The hydrogel pore size influences stem cell migration, proliferation, and differentiation for tissue regeneration. Scanning electron microscope (SEM) images were acquired to characterize the microstructure of freeze-dried Suc-CS / Ald-HA hydrogels prepared at different polymer concentrations (10, 20, 30, 40, and 50 mg / mL) after being incubated at 37 °C for 7 days (FIG. 5). The freeze-dried samples exhibited a continuous and porous structure (FIG. 5), with pore diameters ranging from 8-35 pm depending on the polysaccharide content or concentration of the hydrogels (FIG. 6). Notably, the 8 pm pore size is sufficiently large to facilitate the diffusion of nutrients and metabolites for stem cell culture without hindrance. On the other hand, the 0.2 pm pore size acts as a barrier to the diffusion of large nutrients such as 70 kDa dextran. Consequently, all the Suc-CS / Ald- HA hydrogels described herein ensure the unimpeded diffusion of nutrition and metabolites for stem cell proliferation.

[0083] The pore size and stability of Suc-CS / Ald-HA hydrogel were influenced by the polymer concentration and evolved over the soaking period. On day 0, the pore size inversely correlated with the polymer concentration, as it diminished when the concentration increased from 10 mg / mL to 30 mg / mL. Beyond this point, raising the concentration to 50 mg / mL did not further decrease the pore size significantly, suggesting a threshold in the density of the network that governs pore structure. In terms of homogeneity, the hydrogel at higher concentrations, specifically at 40 and 50 mg / mL, exhibited inhomogeneous porous structure. This inhomogeneity is likely resulting from the high viscosity of the polymer solution, which hindered thorough mixing during gelation. In contrast, at the lower concentration of 10 mg / mL, the hydrogel presented a loosely formed network with irregularly ordered pores, likely due to insufficient crosslinking between -CHO and -NH2 groups. This hydrogel transformed into a fibric structure after one day of incubation and completely dissolved by day 3.

[0084] At a concentration of 20 mg / mL, the pore size decreased from day 0 to day 1, followed by an increase from day 1 to day 7, indicative of a phenomenon characterized by hydrogel contraction followed by swelling. This sequence can be attributed to the initial release of loosely bound water molecules from the hydrogel network upon immersion, resulting in contraction or shrinkage along with the structuralrearrangements within the polymer network. Subsequently, this initial dehydration phase is typically followed by a subsequent swelling phase as the hydrogel absorbs water from its surroundings, possibly driven by osmotic pressure between the polymer chains and water molecules. Consequently, the combined effects of water expulsion and subsequent rehydration contribute to the observed sequence of shrinkage followed by swelling in hydrogel materials. For the concentrations of 30 and 40 mg / mL, the pore size showed a slight enlargement after soaking in DI, likely due to hydrogel swelling. Conversely, at the high concentration of 50 mg / mL, the pore size remained relatively consistent throughout incubation, likely attributable to the higher crosslinking density within the hydrogel structure. These results further demonstrate that the degradation rate of hydrogels of different concentrations varies. The hydrogel with the lowest concentration, 10 mg / mL, degraded the fastest and disappeared completely within 3 days. The 20 mg / mL hydrogel, on the other hand, degraded slowly between day 1 and day 7, while there was no significant degradation of hydrogels with higher concentrations (30, 40, and 50 mg / mL) from day 1 to day 7.

[0085] Biocompatibility of Suc-CS / Ald-HA hydrogels

[0086] The biocompatibility of injectable polymer hydrogels influences the success of the therapeutic intervention by ensuring a harmonious interaction between the hydrogel matrix and the stem cells within the complex biological milieu. Also, healthy hADSCs can secret exosomes to regulate the surrounding tissue microenvironment therapeutically, and they can differentiate into chondrocytes providing an additional cell source to repair cartilage, potentially contributing to the KOA regeneration. As described herein, hADSCs were encapsulated in the hydrogels with different polymer concentrations. To assess the morphology and viability of encapsulated hADSCs, they were stained with live / dead dyes, Calcein-AM and propidium iodide (PI), and visualized using a fluorescence microscope (Olympus 1X81, Olympus Life Science, Waltham, MA, USA). The morphological changes of hADSCs encapsulated in the hydrogels of different polymer concentrations on day 0, 1, 3, 5, and 7 was evaluated. On day 0, at a concentration of 10 mg / mL, some of the hADSCs adhered to the bottom of the cell culture dish, while the remaining cells exhibited a spherical morphology. At concentrations of 20, 30, 40, and 50 mg / mL, the hADSCs maintained their spherical shape, indicating that the hydrogels could potentially serve as a 3D scaffold for supporting cell growth. On day 1, even at a concentration of 10 mg / mL, the cells were able to completely attach and grow on the wellsof microplates (2D culture). This was likely due to the weak and fibrous structure of the hydrogel at this concentration, as seen in FIGS. 4A and 5. Thereafter all cells started 2D proliferation on the well bottom on day 3, as the hydrogel had dissolved completely on day 3. At concentrations of 20, 30, 40, and 50 mg / mL, the hADSCs remained spherical for the entire 7-days period.

[0087] The results of LIVE / DEAD staining of hADSCs encapsulated in hydrogels of different polymer concentrations on days 1, 3, 5, 7, and 14 were determined to evaluate cell viability. On day 1, nearly all hADSCs in all hydrogels remained alive, as indicated by their Calcein-AM staining. hADSCs were shown to be encapsulated homogeneously in the hydrogels with concentrations of 20 and 30 mg / mL. However, cells were not evenly distributed in hydrogels with concentrations of 40 and 50 mg / mL due to the rapid gelation and high viscosity of the hydrogel. At the lower concentration of 10 mg / mL, the hydrogel degraded completely within 3 days, and the cells started to attach and grow on the 2D well plates. These cells adhered and proliferated well on 2D surfaces, maintaining high viability for 14 days. In the hydrogels with concentration of 20 mg / mL and 30 mg / mL, encapsulated cells started to proliferate with high viability and formed increasingly larger cell spheroids over time (FIGS. 7A and 7B). Notably, through this timeframe, nearly all cells encapsulated in 20 mg / mL and 30 mg / mL hydrogel were viable. In the hydrogels with concentrations of 40 and 50 mg / mL, cell viability is low and dead cells increased gradually over time. There is no significant change in the size of cell spheroids over time (FIGS. 7C and 7D), indicating the proliferation of cells was inhibited by the hydrogels at these concentrations. The high crosslinking density and high modulus of these high-polymer concentration scaffolds may hinder the diffusion of nutrients and metabolites causing cell apoptosis.

[0088] Cells encapsulated in hydrogels prepared at 20 and 30 mg / mL, which have a low modulus close to 1000 Pa, exhibited higher viability than those prepared at higher concentrations. This observation aligns with previous studies demonstrating that hydrogels with a softer stiffness of about 1000 Pa tend to preserve stem cell viability, proliferation, and sternness. Furthermore, softer hydrogels can help reduce cell death after transplantation, thus enhancing the therapeutic effects of stem cell injections at the intended location. Representative Suc-CS / Ald-HA hydrogels prepared at 20 mg / mL and 30 mg / mL were demonstrated to be suited for the delivery, long-term proliferation, and maintenance of stem cells.

[0089] Expression of genes relevant to cartilage regeneration

[0090] The hydrogels described herein, cartilage regeneration is facilitated by chondrogenesis differentiation of hADSCs encapsulated in a hydrogel. An examination of hADSC differentiation into chondrocytes at the gene expression level was examined. The expression of specific genes, including ACAN (Aggrecan), SOX9 (SRY-box transcription factor 9), and COL2A1 (Collagen Type II Alpha 1 chain), was examined. ACAN is maintains cartilage integrity, as it encodes the aggrecan protein that forms the foundation of the cartilage extracellular matrix, providing hydration and resistance. SOX9, a transcription factor, regulates chondrocyte differentiation and cartilage formation by controlling the expression of genes like ACAN and COL2A1. Type II collagen (Col-II), as a major component of cartilage matrix, provides structural stability and strength.

[0091] The relative expression of these genes in various culture conditions was evaluated. Cells were cultured on 2D tissue culture polystyrene (TCPS), a conventional platform, and in 3D hydrogels at two concentrations: 20 nig / mL and 30 nig / mL in chondrogenic differentiation medium. Notably, the 3D hydrogel environment closely mimics the native cartilage microenvironment. FIGS. 8A-8C show that cells cultured in differentiation medium exhibited significantly elevated expression levels of ACAN, SOX9, and COL2A1 on both day 7 and day 14 when compared to the cell cultured in the complete medium. Also, the results demonstrated the capability of the 3D hydrogel platform in enhancing gene expression associated with chondrogenic differentiation compared to the 2D TCPS surface on both day 7 and day 14. Interestingly, no significant difference in gene expression (ACAN and SOX9) was observed between cells encapsulated in the two different hydrogel concentrations on day 7. However, on day 14, the 3D hydrogel formulated at a concentration of 30 mg / mL displayed a remarkable ability to enhance gene expression and chondrogenic differentiation compared to the 20 mg / mL hydrogel. This distinctive response on day 14 between the two hydrogel concentrations could be attributed to the mechanical properties and scaffold architecture of the hydrogels. The 30 mg / mL hydrogel might offer improved mechanical support and structural cues, facilitate cellular organization and interaction, and ultimately enhance chondrogenic differentiation and gene expression. Additionally, the higher concentration might provide a more conducive microenvironment for chondrocyte maturation, influencing cellular signaling pathways and extracellular matrix deposition. These findings demonstrated the potential of 3D hydrogel systems, especially at a 30 mg / mL concentration, as a promising approach for effectivecartilage regeneration, highlighting the significance of microenvironmental factors in guiding chondrogenic differentiation and tissue development. Also, results showed that hADSCs encapsulated within the 3D hydrogel exhibited elevated expression levels of ACAN, SOX9, and C0L2A1 in comparison to cells cultured on the two-dimensional (2D) surface, even when the cells are maintained in a normal stem cell culture medium rather than a chondrogenic differentiation medium. This further indicates that the hydrogels described herein can stimulate chondrogenic differentiation.

[0092] The gene expression fold-changes observed in a 3D hydrogel system, particularly at the 30 mg / mL concentration, are comparable to or greater than those reported in other hADSC-based scaffolds for cartilage regeneration. For instance, about 24-fold, about 177-fold, and about 533-fold increases in SOX9, ACAN, and COL2A1 expression, respectively, after 21 days in an atelocollagen hydrogel have been reported (Kim et al., Scientific Reports, 2020, 10, 10678). Similarly, it has been observed that COL2A1 upregulation in the range of 102-103-fold in hADSCs cultured on a silk fibroin scaffold combined with platelet-rich plasma (Rosadi et al., Stem Cell Research & Therapy, 2019, 10, 369). As described herein, ACAN, SOX9, and COL2A1 expression levels exceeded 1000-fold on day 14 within the 30 mg / mL hydrogel group, indicating a strong chondrogenic differentiation response. These results not only align with but in some cases surpass previously reported outcomes, highlighting the effectiveness of the hydrogel system described herein in supporting cartilage-specific gene expression.

[0093] Expression of proteins relevant to cartilage regeneration

[0094] Immunofluorescence staining was used to assess the expression of chondrogenic differentiation-associated proteins in hADSCs encapsulated in hydrogels. Specifically, the evaluation focused on ACAN and SOX9 expression, which was visualized using green and red fluorescence, respectively. Concurrently, cell nuclei were counterstained with DAPI in blue. ACAN is a key extracellular matrix protein that plays a role in cartilage formation and maintenance. Its presence is indicative of a mature and functional chondrocyte phenotype, underscoring it as a marker for chondrogenic differentiation. SOX9, on the other hand, is a transcription factor responsible for the regulation of genes involved in cartilage development and maintenance. The expression of SOX9 serves an indicator of chondrogenic commitment, as it orchestrates the differentiation process and ensures the proper formation of cartilaginous tissues.

[0095] The results clearly demonstrate that hADSCs cultured within 3D hydrogels (at concentrations of 20 and 30 mg / mL) (3D-d) and on a 2D surface in a chondrogenic differentiation medium (2D-d) consistently exhibited robust expression of ACAN and SOX9. In contrast, hADSCs cultured on the 2D surface in a regular cell culture medium (2D) displayed negligible expression of ACAN and SOX9. While hADSCs cultured in both 3D-d and 2D-d environments exhibited expression of chondrogenic differentiation-associated proteins (ACAN and SOX9), distinct advantages are offered by 3D hydrogels over 2D surfaces. These advantages stem from the hydrogels’ ability to closely mimic the native tissue microenvironment, providing spatial cues and mechanical support useful for cellular organization and function. Unlike 2D surfaces, which limit cellcell and cell-matrix interactions, 3D hydrogels facilitate multidirectional interactions between cells and the surrounding matrix, thereby promoting physiological responses and offering a more accurate representation of in vivo conditions. Additionally, the porous structure of hydrogels enables enhanced diffusion of nutrients and signaling molecules, thereby further enhancing cellular behavior and promoting the development of complex tissues. Consequently, these findings suggest that the 3D hydrogel environment can effectively support the chondrogenic differentiation of hADSCs, as evidenced by the notable expression of ACAN and SOX9, markers associated with commitment to the chondrogenic lineage.

[0096] In parallel with the observed gene and protein expression patterns, the degradation behavior of the hydrogels was evaluated to better understand their suitability for supporting dynamic tissue remodeling. The 20 mg / mL and 30 mg / mL hydrogels both exhibited an initial increase in weight due to swelling, followed by a gradual decrease in mass over a 14-day period. This controlled and sustained degradation profile is useful for injectable hydrogels, as it provides temporary structural support during the early phase of chondrogenic differentiation, while allowing sufficient space for extracellular matrix deposition as new tissue forms. Notably, this degradation timeline coincides with the period of robust ACAN, SOX9, and COL2A1 gene upregulation and strong protein expression, particularly in the 30 mg / mL group, indicating that the hydrogel degrades at a rate conducive to cell-mediated matrix production and cartilage regeneration. These findings suggest that the degradation behavior of the hydrogels described herein not only support but may actively facilitate the transition from a synthetic scaffold to a tissue- integrated construct.

[0097] In one aspect, the present disclosure provides a cell-compatible, biodegradable, injectable, and self-healing polysaccharides-based hydrogel (Suc-CS / Ald- HA) in situ through simultaneous injection of two polymer solutions. The resulting hydrogel not only provided a favorable environment for both delivering and maintaining hADSCs with high viability, demonstrating cartilage regeneration. The gelation occurred upon mixing the two solutions to form a homogeneous, highly porous structure. By changing the concentrations of the polymer solutions, the hydrogel is tailored to achieve the desired mechanical properties, pore size, and degradation rate, ensuring the sustained viability of encapsulated hADSCs. The encapsulated hADSCs remained highly viable even after 14-days incubation, demonstrating their potential for supporting cartilage regeneration over an extended duration. Moreover, the hADSCs encapsulated within the 3D hydrogels exhibited elevated gene expression levels associated with chondrogenic differentiation. Notably, the 3D hydrogels did not hinder the expression of proteins related to chondrogenic differentiation in the encapsulated hADSCs. This injectable stem cellladen hydrogel holds promise not only for cartilage regeneration but also for potential applications in other cell-based therapies.

[0098] Tissue Adhesive Chitosan / Hyaluronic Acid-Based Hydrogels

[0099] In another aspect, the disclosure provides chitosan / hyaluronic acid-based hydrogels that further include a tissue adhesive component that imparts tissue adhesiveness to the hydrogel.

[0100] Effective tissue adhesion is vital for cell-laden hydrogels in knee cartilage regeneration, ensuring encapsulated cells remain at targeted sites for sustained therapeutic impact. Despite chitosan’s natural mucoadhesive properties, chitosan-based injectable hydrogels often exhibit poor tissue adhesion, primarily due to their higher water content, which hampers adhesion to tissue. Chemical modifications have thus been explored to enhance the adhesive properties of these hydrogels. Thiolated chitosan demonstrates strong adhesion to biological surfaces, but this adhesion is transient in vivo due to reversible disulfide bridges. In contrast, catechol groups offer longer-lasting adhesion. Catechol, which occurs naturally as a functional group in L-3,4-dihydroxyphenylalanine (L-DOPA), is readily available in derivative forms and is known to engage in strong hydrogen bonding and coordination interactions with biological substrates. Catechol oxidizes to quinone, enabling covalent bonding with tissue protein amine and thiol groups, collectively ensuring durable adhesion. Beyond adhesion, DOPA offers anti-inflammatoryand antioxidant benefits and contributes to neurotransmitter synthesis, enhancing its therapeutic value.

[0101] The preparation, characterization, and use of a representative chitosan / hyaluronic acid-based hydrogel (Ald-HA / Dopa-HBCS hydrogel) is described below.

[0102] The present disclosure provides a biocompatible, injectable, tissueadhesive polysaccharide hydrogel designed to sustain human adipose-derived stem cells (hADSCs) at cartilage defect sites within the joint cavity, promoting cartilage regeneration (FIG. 9). This hydrogel was engineered using a dual physical and chemical crosslinking strategy. To enhance chitosan’s solubility and adhesion, hydroxybutyl and catechol groups were introduced to its backbone, yielding catechol-functionalized hydroxybutyl chitosan (Dopa-HBCS). The amino groups of Dopa-HBCS were leveraged to form a chemically crosslinked network with modified hyaluronic acid (HA) via a Schiff base reaction. HA is a linear polysaccharide composed of alternating units of N-acetyl-D-glucosamine and D- glucuronic acid and serves as an extracellular matrix component with excellent biodegradability and biocompatibility. Nevertheless, the clinical utility of HA in cartilage repair and related applications has been hindered by its rapid enzymatic degradation and inherently high viscosity. As described herein, HA's carbon-carbon bonds were cleaved through oxidation to generate aldehyde groups, forming aldehyde HA (Ald-HA). These aldehyde groups react with Dopa-HBCS amino groups via a Schiff base reaction, creating a robust covalent network and yielding an injectable, tissue-adhesive double -network hydrogel (e.g., Ald-HA / Dopa-HBCS). The hydrogel’s physicochemical properties (e.g., pore size, degradation rate, cellular adhesivity, and rheological behavior) are described herein. hADSCs were encapsulated during gelation, and their morphology, viability, and activity were evaluated at various culture time points.

[0103] HBCS polymer synthesis and characterization

[0104] To enhance the solubility of chitosan under physiological conditions, the polymer was chemically modified via hydroxybutylation. Chitosan was dissolved into DI containing KOH and urea, and two freeze-thaw cycles were performed to obtain a transparent and homogeneous chitosan solution. Potassium ions (K+) and urea disrupt the native hydrogen bonding and hydrophobic interactions within the chitosan matrix, thereby enhancing its solubility and promoting a more homogeneous reaction environment. After the chitosan was completely dissolved, 1,2-epoxybutane was directly added into thechitosan solution. The synthesis of HBCS involved the nucleophilic ring -opening of 1,2- epoxybutane by hydroxyl groups on the chitosan backbone (FIG. 10 A). Under alkaline conditions, the hydroxyl groups on the chitosan backbone exhibit greater nucleophilicity than the amino groups, making them the preferred reactive sites for the ring-opening reaction with 1,2-epoxybutane. In the proton nuclear magnetic resonance (1H NMR) spectrum of chitosan, HBCS, and Dopa-HBCS, a distinct peak at 0.9 ppm, observed in the 'H NMR spectra of both HBCS and Dopa-HBCS but absent in native chitosan, was attributed to the methyl protons (-CH3) of the hydroxybutyl substituents. Successful incorporation of hydroxybutyl groups into the chitosan backbone was further supported by FT-IR analysis, which revealed new absorption bands at approximately 1570 cm1and 2920 cm1, attributable to the asymmetric stretching of methylene (-CH2) and methyl (- CHs) groups. To eliminate the interference of D2O peaks with any characteristic peaks and obtain the degree of substitution (DS) of hydroxybutyl groups, 20% DCI / D2O was used as the1H NMR solvent. The DS was evaluated to be 65.3% by comparing the integral of the peaks of anomeric protons (Hl, 4.5-4.8 ppm) and the integral of the peaks of the methyl protons (-CH3, 0.9 ppm). The DS value suggested that the substitution reaction predominantly took place at the hydroxyl groups situated in the C6 position.

[0105] Dopa-HBCS was synthesized through carbodiimide-mediated coupling in an acidic environment, where EDC activated the carboxyl groups of 3,4- dihydroxyphenylacetic acid, and NHS stabilized the intermediate, facilitating conjugation to the primary amines on HBCS (FIG. 10A). The appearance of a new absorption band at 1725 cm1in the FT-IR spectrum, absent in both native chitosan and HBCS, corresponds to C=O stretching vibrations of amide and ester linkages, confirming the successful conjugation of catechol groups to the polymer backbone. Also, newly observed peaks in the 6-7 ppm range of the 'H NMR spectrum confirmed the presence of catechol protons. The DS of catechol groups was evaluated to be 14.9 %. To confirm that the hydroxyl groups on the catechol moiety remained in their reduced form and were not oxidized to quinones, ultraviolet-visible absorption (UV-Vis) spectra of Dopa-HBCS were measured. The presence of an absorption band at 280 nm verified successful catechol conjugation, while the lack of a corresponding peak near 400 nm suggested that the phenolic hydroxyl groups remained unoxidized.

[0106] To synthesize Ald-HA, HA was reacted with sodium periodate. Sodium periodate selectively cleaves vicinal diols within the sugar ring through oxidative cleavage,leading to the formation of dialdehyde derivatives. The addition of aldehyde groups to HA was achieved through this process. The reaction scheme and resulting chemical structure of Ald-HA are illustrated in FIG. 10B. The presence of aldehyde groups in the Ald-HA was evidenced by a distinct resonance at approximately 5 ppm in the1H NMR spectrum. The oxidation percentage of HA was 40% based on aldehyde quantification using tert-butyl carbazate derivatization. This value was calculated by comparing the integration of the tert-butyl proton signal at 8 = 1.38 ppm with that of the acetamide protons of HA at 8 = 1.90 ppm in the1H NMR spectrum.

[0107] Ald-HA / Dopa-HBCS hydrogel preparation and characterization

[0108] Injectable hydrogels can be fabricated using two strategies. Firstly, in situ gelation of liquid polymers post-injection involves the combination of precursor solutions that undergo gelation once injected into the target site. This method allows for a controlled and localized formation of the hydrogel at the site of implantation or injection. Alternatively, the application of shear-thinning polymer systems represents a second approach, wherein the hydrogel precursor exhibits a temporary reduction in viscosity under shear stress during injection, facilitating its smooth administration. Once injected, the polymer system recovers its original viscosity, enabling in situ gelation and stable hydrogel formation at the target site. Here, the hydrogel was created by gently mixing the solutions of modified polysaccharides: Ald-HA and Dopa-HBCS using the first strategy. A medicalgrade dual-barrel syringe fitted with a 21 -gauge needle was utilized to uniformly mix polymer solutions during the injection which allows the in situ gelation post-injection.

[0109] FIG. 11 shows the Ald-HA / Dopa-HBCS hydrogel’s chemical structure, where aldehyde groups of Ald-HA and amino groups of Dopa-HBCS form a covalent imine network via a Schiff base reaction, supplemented by hydrophobic and electrostatic interactions between hydroxybutyl groups and charged chitosan and HA. Initial mixing of Ald-HA with HBCS or Dopa-HBCS in deionized water without CaCh produced a white precipitate, likely from polyelectrolyte complexes of cationic chitosan and anionic HA, increasing heterogeneity. Adding 100 mM CaCb to the solutions prevented precipitation by shielding electrostatic interactions, yielding a transparent, homogeneous hydrogel. CaCb minimizes free calcium ions, with gradual ion exchange in physiological conditions reducing risks of calcium phosphate precipitation and cartilage calcification, supported by low cytotoxicity, as excess Ca2+could disrupt cellular homeostasis and induce cell death.

[0110] Results indicate that both HBCS and Dopa-HBCS solutions remain liquid but form hydrogels upon mixing with Ald-HA across all concentrations tested (5, 10, 20, and 30 mg / mL). Higher polymer concentrations were found to accelerate gelation. Gelation time decreased with increasing concentration, dropping from 280 seconds at 5 mg / mL to 5 seconds at 30 mg / mL, demonstrating that polymer concentration significantly influences gelation kinetics. This accelerated gelation at higher concentrations stems from enhanced polymer-polymer interactions, which reduce the time needed to form a stable gel network. The increased likelihood of crosslinking events, driven by greater polymer chain density, accounts for the observed decrease in gelation time. The gelation time at 10 mg / mL (about 30-60 s) is suitable for clinical syringe injection, allowing sufficient working time for in situ application while minimizing premature leakage at the target site.

[0111] The mechanical properties of hydrogels play a role in stem cell proliferation and phenotype. Therefore, rheological analysis was conducted to evaluate hydrogels prepared at different polymer concentrations and with or without CaCL (FIGS. 12A-12D). The mechanical strength of the Ald-HA / Dopa-HBCS hydrogels were examined by stress sweeping testing at 37 °C. Hydrogels were prepared by mixing two respective polymer solutions, each containing equal volumes and concentrations, and were tested to determine the hydrogel with desired mechanical properties. The storage modulus (G’) reflects the elastic component of a material’s viscoelastic response, corresponding to its solid-like behavior, while the loss modulus (G”) reflects the viscous component, indicative of liquid-like behavior. In crosslinked hydrogels, the dominant elastic network results in G’ exceeding G” (G’ > G”), characteristic of viscoelastic solids. In contrast, uncrosslinked polymer solutions exhibit G” values greater than G’ (G’ < G”), as the absence of a stable network leads to predominantly viscous behavior due to weak or transient interactions among polymer chains.

[0112] FIGS. 12A-12D shows that G’ increased across all hydrogel concentrations with CaCL compared to those without, where unchecked electrostatic interactions caused precipitation. CaCL suppresses these interactions, enabling Schiff base bond formation and yielding a transparent hydrogel, confirming Schiff base-driven gelation. In A5H5 (+) versus A5D5 (+) with CaCL, dopamine conjugation raised G’ at 5 mg / mL but lowered it at 20 and 30 mg / mL, with no notable G’ difference between Ald- HA / HBCS and Ald-HA / Dopa-HBCS at 10 mg / mL. At 5 and 10 mg / mL, abundant dopamine groups promote covalent crosslinking via oxidative polymerization, forming astrong network, while at higher concentrations, increased Schiff base-reactive sites (amino and carbonyl groups) favor Schiff base reactions. For A5D5 (+) hydrogels, G’ and G” remained stable in the linear viscoelastic regime (LVR), with critical shear stress rising from 5 Pa to 200 Pa as concentration increased from 5 to 30 mg / mL, indicating enhanced crosslinking and structural resilience. G’ and G” stability under shear stress highlights the gel’s robust mechanical performance across loads.

[0113] Previous research has indicated that hydrogels possessing a G’ between 500 and 1000 Pa provide a mechanically favorable environment for cartilage regeneration, whereas stiffer hydrogels with G’ approaching 3000 Pa are more appropriate for the regeneration of fibrous tissues. Thus, the 10 mg / mL Ald-HA / Dopa-HBCS hydrogel was selected for subsequent investigation base on its optimal mechanical properties and gelation time, and the Ald-HA / HBCS hydrogel was used as reference material.

[0114] Next, the effect of volume mixing ratios of polymer solutions on the mechanical properties and gelation time of formed hydrogels was evaluated. FIG. 12E shows the rheological analysis of Ald-HA / Dopa-HBCS hydrogels at different volume mixing ratios at 10 mg / mL in DI containing 100 mM CaCL. For all Ald-HA / Dopa-HBCS hydrogels, those prepared at the same volume mixing ratio (A5H5 and A5D5) showed the highest storage modulus (G’) around 400 Pa. In contrast, hydrogels prepared at a low volume mixing ratio of Ald-HA (A3H7 and A3D7) showed relatively lower storage modulus (G’) around 150 Pa. The hydrogels prepared at higher volume mixing ratio of Ald-HA (A7H3 and A7D3) showed a storage modulus (G’) around 50 Pa. These hydrogels were too weak and completely dissociated in DPBS after being incubated at 37 °C for 24 h, unsuitable as injectable gels. These results indicated that a high degree of crosslinking between polymer chains could be obtained when the hydrogels were prepared at the same volume mixing ratio of Ald-HA and Dopa-HBCS. The mechanical strength of hydrogels affects cell proliferation behavior and the therapeutic efficacy for tissue regeneration. Thus, A5D5 hydrogel of 10 mg / mL was selected as the optimized hydrogel for further cell experiments, and A5H5, A3H7 and A3D7 hydrogels were selected as reference standard.

[0115] The hydrogel’s capability to retain cells post-injection at defect sites plays a role in promoting the effectiveness of cell-based therapeutic interventions. To evaluate this, two polymer solutions, Ald-HA and Dopa-HBCS were blended through a double barrel syringe equipped with a G21 needle and injected onto the measurement area of a rheometer (Anton Paar MCR 92, USA). The G’ and G” were recorded over time under atime sweeping mode to study the variation of mechanical strength of the hydrogel in realtime, and a 1 Hz frequency and a 1% strain were used. The measurement started immediately after the injection of the hydrogel at 37 °C to mimic physiological temperature. As mentioned above, the G’ of viscoelastic solid hydrogel is higher than its G”, whereas the G’ of the viscoelastic liquid solution is lower than its G”. Therefore, each polymer solution as the hydrogel precursor would exhibit the viscoelastic liquid solution behavior (G’ < G”) before gelation, whereas the hydrogel would exhibit the viscoelastic solid hydrogel behavior (G’ > G”).

[0116] FIG. 12F shows that the storage modulus was greater than loss modulus (G’ > G”) for all the samples, continuing the successful gel formation. In addition, crossover points between G’ and G” was not detected, suggesting that all the polymer mixtures underwent a rapid transition from solution to hydrogel right after injection within 1 min. The hydrogel was created by a combined physical and chemical crosslinking approach: (1) the physical crosslinking induced by the intermolecular hydrophobic interaction at 37 °C, and (2) the chemical crosslinking between aldehyde and amino groups through Schiff base linkage. Within 10 min, the double crosslinking enabled the storage modulus to increase to the range of 150-200 Pa, which has been previously reported to effectively encapsulate the embedded cells within the 3D structure, preventing cell leakage after injection and improving the efficiency of cell therapy. Additionally, the A5D5 hydrogel at 10 mg / mL in complete stem cell culture medium exhibited rapid self-healing behavior, with more than 98% of G’ recovered within 3 minutes. The self-healing property of the hydrogel in cell-laden hydrogel-based therapy for knee repair ensures the sustained structural integrity and functionality of the hydrogel, thereby promoting prolonged support for cell growth and tissue regeneration; additionally, this self-healing capacity may enhance clinical performance by maintaining adhesion and mechanical stability under joint motion and physiological loading conditions.

[0117] SEM imaging was performed to investigate the microarchitecture of the freeze-dried hydrogels: A5D5, A3D7, A5H5, and A3H7 (FIG. 13 A). The hydrogel pore size was determined by analyzing SEM images with ImageJ software (FIG. 13B). All the freeze-dried hydrogels exhibited an interconnected porous structure, with pore sizes ranging from 15 to 35 pm. The hydrogel prepared at the same volume mixing ratio of A5D5 and A5H5 showed a more uniform porous structure than the hydrogel prepared at a low volume mixing ratio of Ald-HA (A3D7 and A3H7). The A5D5 hydrogel containingcatechol groups showed a smaller pore size (16.2 pm) than the A5H5 hydrogel without catechol groups (28.1 pm), likely due to the additional crosslinking formed between adjacent catechol groups through covalent bonding and TI- interactions. Generally, the 8 pm pore size is large enough to facilitate the proper diffusion of nutrition and metabolites, while the 0.2 pm pore size restricts the diffusion of large nutrients (e.g., 70 kDa dextran). The observed pore size (15-35 pm) supports efficient nutrient diffusion and may promote cell retention and clustering, which are favorable for stem cell viability and chondrogenic differentiation. Thus, all the Ald-HA / Dopa-HBCS hydrogels described herein provide full diffusion of nutrition and metabolites useful for stem cell proliferation. Also, the pore size can be further modulated by adjusting polymer concentration, freezing conditions, and crosslinking density to suit different applications. Moreover, the result of storage modulus and pore size collectively suggested that A5D5 hydrogels are effective in retaining stem cells within the hydrogel matrix while facilitating the diffusion of essential nutrition and metabolites. Although some pore sizes (15-35 pm) were smaller than the average diameter of individual stem cells (about 20 pm), the hydrated and viscoelastic nature of the hydrogel permits cell deformation, local remodeling, and adhesion within the matrix.

[0118] Hydrogel degradation was assessed by measuring changes in dry weight over time, with lyophilized samples collected at designated intervals and compared to the initial dry weight of freshly prepared hydrogels. As shown in FIG. 13C, the weight change of A5D5 hydrogels was comparable with that of A5H5, and the weight change of A3D7 hydrogels was comparable with that of A3H7. The degradation rates of A5D5 and A5H5 were much slower than those of A3D7 and A3H7 because the A5D5 and A5H5 would have higher crosslinking densities leading to higher stability than A3D7 and A3H7. A5D5 hydrogels showed a slow degradation rate and more than 80% of the hydrogel remained after 14 days. A long-term follow-up experiment showed that more than 50% of the A5D5 hydrogel remained after 40 days. Notably, previous research reported that a thermoresponsive hydrogel made of solely Dopa-HBCS showed fast weight loss, completely disappearing within 14 days. This suggests the significant enhancement in stability achieved by incorporating Ald-HA through the Schiff base reaction, making the hydrogels described herein suitable for delivering stem cell and regenerating cartilage.

[0119] Biocompatibility of Ald-HA / Dopa-HBCS hydrogels

[0120] The biocompatibility of hydrogels plays a role in determining the functionality of encapsulated human adipose-derived stem cells (hADSCs) and,consequently, their therapeutic efficacy. To evaluate biocompatibility, polymers were dissolved in complete mesenchymal stem cell culture medium, and cell-laden hydrogels (A5D5, A3D7, A5H5, A3H7) containing equal numbers of hADSCs were prepared in 96- well plates. Cell viability of encapsulated hADSCs was examined by fluorescent labeling with Calcein-AM for live cells and propidium iodide for dead cells and visualized using fluorescence microscopy. Calcein-AM penetrates live cells, where it is metabolized into calcein, emitting green fluorescence, while propidium iodide, impermeable to intact cell membranes, binds to DNA in dead cells, emitting red fluorescence upon loss of membrane integrity.

[0121] hADSC viability in hydrogels with varying volume mixing ratios, with or without catechol groups, was evaluated on days 1, 3, 5, 7, and 14 post-encapsulation. From day 1 to day 5, hADSCs were uniformly encapsulated within all hydrogels, displaying a spherical morphology, indicating that these hydrogels effectively served as 3D scaffolds, maintaining high cell viability for at least 5 days. By day 7, some cells were released from the hydrogels, likely due to hydrogel degradation, and these released cells proliferated in a monolayer fashion. Fewer cells were released from A5D5 and A5H5 hydrogels compared to A3D7 and A3H7, reflecting the slower degradation rates of A5D5 and A5H5. From day 1 to day 14, cell numbers in all hydrogels increased over time due to proliferation. Remarkably, throughout this period, nearly all cells remained viable (green fluorescence), with minimal dead cells (red fluorescence) observed, confirming the biocompatibility of the hydrogel system. The Live / Dead assay was conducted to evaluate hydrogel biocompatibility, with the cell density and volume selected for imaging consistency in vitro rather than to replicate in vivo conditions. The injected hydrogel volume and cell density can be readily adjusted based on specific experimental or clinical requirements.

[0122] Cell morphologies within the cell-laden hydrogels (A5D5, A3D7, A5H5, A3H7) on day 14 were determined. The images showed that cells formed spheroids, with those in A3D7 and A3H7 appearing slightly larger than those in A5D5 and A5H5. This size difference likely results from the faster degradation of A3D7 and A3H7, which may have created a more permissive environment for spheroid growth compared to the more stable A5D5 and A5H5 hydrogels. Also, stem cell viability remained unaffected by the injection process using a dual-barrel syringe, preserving their integrity and functionality. This precise delivery method highlights its suitability for maintaining stem cell viability, a factor for effective therapeutic applications.

[0123] The injectable hydrogel’s primary function is to support the proliferation of encapsulated stem cells, creating an optimal environment for advancing stem cell-based knee cartilage repair. FIG. 14 shows quantitative analysis of cell proliferation based on metabolic activity measured by the Alamar Blue assay. On day 1, fluorescence intensities for A5H5 and A3H7 were similar. Over time, A3H7 exhibited a steady increase in fluorescence intensity from day 1 to day 14, reflecting sustained cell proliferation within the hydrogel. Conversely, A5H5 showed minimal change in fluorescence intensity until day 5, followed by a marked rise from day 5 to day 14, confirming the sustained viability of hADSCs within the hydrogel. On day 1, fluorescence intensities for A5D5 and A3D7 were lower than those for A5H5 and A3H7, suggesting that catechol group conjugation initially suppressed hADSC proliferation. However, by day 7, cell numbers in A5D5 and A3D7 reached levels comparable to those in A5H5 and A3H7, likely due to catechol groups enhancing hADSC adhesion to the hydrogel matrix, thereby promoting cell growth.

[0124] Moreover, hADSC proliferation was greater in hydrogels with a lower Ald-HA volume mixing ratio (A3D7, A3H7) compared to those with a higher ratio (A5D5, A5H5). This difference may stem from the elevated aldehyde content in A5D5 and A5H5, which could impede cell growth. Aldehyde groups may disrupt extracellular matrix interactions, induce oxidative stress, and suppress proliferation-related gene expression, collectively limiting hADSC expansion in crosslinked hydrogels. These findings raise concern about potential cytotoxic effects associated with unreacted aldehyde groups at higher concentrations. Although free aldehyde groups can potentially be cytotoxic, this concern is mitigated by rapid Schiff-base crosslinking during gelation, which effectively consumes reactive groups and stabilizes the hydrogel environment for encapsulated stem cells. As a result of this rapid reaction and the hydrogel’s overall biocompatible design, both volume mixing ratios of these hydrogels exhibited favorable biocompatibility. Although direct quantification of residual aldehyde content was not performed, the efficient gelation, formation of stable porous structures across different volume mixing ratios, and high cell viability over 14 days collectively indicate that most aldehyde groups were effectively consumed during the Schiff-base crosslinking process. In addition, these standard in vitro conditions (complete medium, 5% CO2, 37 °C) were used to evaluate hydrogel biocompatibility and cell viability in a controlled and reproducible manner.

[0125] Tissue adhesive strength of Ald-HA / Dopa-HBCS hydrogels

[0126] The ability of a hydrogel to adhere to native tissue plays a role in stem cell-based tissue regeneration. Such an ability not only enables the gel-cell to construct to remain at the therapeutic site but also provides mechanical support and protection for encapsulated cells. Most of the current hydrogels have demonstrated a limited ability in tissue adhesion which can compromise their effectiveness in vivo. In this design, a bioinspired catechol functional group was incorporated to strengthen interfacial interactions between the hydrogel and surrounding tissue, thereby enhancing adhesive properties and facilitating improved integration with the host environment. The catechol group has shown promising results in preclinical studies and offer significant promise for improving the effectiveness of hydrogel -based therapies in tissue engineering applications.

[0127] One current challenge in cartilage tissue engineering is the limited integration of implanted hydrogel with the native tissue. In the design described herein, the incorporation of catechol groups in Dopa-HBCS enhances tissue adhesion by facilitating multiple types of interactions at the hydrogel-tissue interface, including covalent reactions, noncovalent n-n stacking, and hydrogen bonding. The covalent reactions create strong adhesion capability. Upon oxidation by dissolved oxygen, strong oxidizing agents, or under alkaline conditions, catechol groups are converted into highly reactive o-quinone intermediates. The o-quinone undergoes Michael-type addition or Schiff base formation reactions with lysine or cysteine residues present in the tissue surface, resulting in the formation of covalent interfacial bonds. Also, polyphenol crosslinks will form between o-quinone groups which will contribute to the elastic properties of synthetic adhesive hydrogels (FIG. 15A). Unlike thiol-based adhesives, which degrade under oxidative conditions, catechol groups remain stable, ensuring longterm adhesion. In contrast to fibrin-based adhesives, which require enzymatic activation, catechol-functionalized hydrogels adhere instantly through strong interfacial interactions. This rapid, durable bonding to wet or dynamic tissues makes them useful for biomedical applications requiring reliable and long-lasting adhesion across various substrates.

[0128] To assess the function of catechol groups to tissue adhesive properties, lap shear tests were performed on three hydrogel formulations: A5H5, A5D5, and A5DH5. The A5DH5 hydrogel was prepared by mixing HBCS and Dopa-HBCS at a 5:5 volume ratio, followed by combination with Ald-HA solution at a 5:5 volume ratio (Ald-HA, Dopa- HBCS, and HBCS were mixed at a volume ratio of 5:2.5:2.5, v / v / v). Hydrogels were applied to the overlapping region of two porcine skin samples at room temperature using adouble-barrel syringe. After overnight incubation at 37 °C, adhesion strength was measured using a Shimadzu Autograph AGS-X universal testing machine under a constant shear rate of 1.3 mm / min until separation occurred. All samples exhibited adhesive failure.

[0129] As shown in FIGS. 15B and 15C, A5H5 (Ald-HA and HBCS) displayed minimal adhesive strength (2.2 kPa). In contrast, A5DH5 (Ald-HA with a Dopa-HBCS and HBCS mixture) showed a notable increase to 2.8 kPa. The highest adhesive strength, approximately 4.1 kPa, was observed with A5D5 (Ald-HA and Dopa-HBCS), which contains the greatest catechol group content. With a catechol grafting ratio of 14.9% in Dopa-HBCS, adhesion strength can be further tuned by adjusting this ratio. The adhesive strength of A5D5 over A5DH5 reflects its higher catechol concentration, confirming that catechol groups significantly enhance tissue adhesion, with strength increasing alongside catechol content. Additionally, adhesive interfaces exhibit greater rigidity, reducing slippage or deformation. Consequently, under external stress, these materials show minimal deformation, resulting in lower maximum strain in hydrogels with higher adhesive strength. The adhesive strength of A5D5 aligns with that of other catechol-based hydrogels and meets the requirements for tissue engineering scaffolds.

[0130] Overall, A5D5 proved to be an injectable, biocompatible, and tissueadhesive polysaccharide hydrogel, likely capable of retaining hADSCs at cartilage defect sites to facilitate cartilage regeneration via cell therapy. Schiff-base crosslinking may further bolster tissue adhesion by forming covalent bonds between Ald-HA aldehyde groups and tissue proteins, enhancing adhesion strength and hydrogel retention. As described herein, Ald-HA aldehyde groups primarily drive hydrogel formation, with crosslinking optimized by adjusting the Dopa-HBCS to Ald-HA ratio. Nonetheless, some aldehyde groups inevitably interact with tissue, further supporting adhesion.

[0131] Genes expression related to cartilage regeneration

[0132] The expression of specific genes aggrecan (ACAN), SRY-box transcription factor 9 (SOX9), and Collagen Type II Alpha 1 chain (COL2A1) plays a role in regeneration of cartilage. ACAN encodes aggrecan, a major proteoglycan in the extracellular matrix of cartilage for matrix hydration and resistance to compressive forces. SOX9 functions as a master transcription factor regulating chondrocyte lineage commitment and differentiation by activating downstream targets, including ACAN and COL2A1. Collagen type II (COL-II), encoded by COL2A1, is a principal structural protein in cartilage that provides tensile strength and mechanical integrity.

[0133] To evaluate the influence of different culture environments on the expression of these genes, hADSCs were cultured either on standard 2D tissue culture polystyrene (TCPS) or within 3D hydrogels in a chondrogenic differentiation medium. The 3D hydrogel environment closely emulates the native cartilage microenvironment. As shown in FIGS. 16A-16C, cells cultured in the 3D hydrogel exhibited significantly higher expression levels of ACAN, SOX9, and COL2A1 on days 7 and 14 compared to those on 2D TCPS. These findings suggest that the 3D hydrogel environment enhances chondrogenic gene expression, likely through improved cell-matrix interactions and a more physiologically relevant mechanical milieu. The enhanced structural mimicry may promote cellular condensation, matrix production, and signaling pathways for chondrocyte maturation. Overall, these results demonstrate the potential of the 3D hydrogel systems as an effective platform for promoting cartilage regeneration, highlighting the use of biomimetic microenvironments in directing stem cell fate and tissue-specific gene expression.

[0134] In one aspect, the present disclosure provides in wm-formed biocompatible, biodegradable, injectable, self-healing, and tissue -adhesive polysaccharide- based hydrogel (e.g., composed of aldehyde-functionalized hyaluronic acid (Ald-HA) and dopamine-modified hydroxybutylchitosan (Dopa-HBCS)). This hydrogel system offers a favorable microenvironment for the minimally invasive encapsulation and sustained support of hADSCs, ensuring high cell viability and exhibiting strong potential for cartilage regeneration. Gelation was rapidly achieved upon mixing the two polymer components, forming a uniform and highly porous network of hydrogels. By adjusting the concentrations and volume mixing ratios of the polymer solutions, the hydrogel’s mechanical strength, gelation kinetics, and degradation profile to support prolonged viability of encapsulated hADSCs was optimized. Post-injection, encapsulated hADSCs exhibited high viability for at least 14 days, confirming the hydrogel’s suitability for cartilage repair applications. The incorporation of catechol groups conferred tissueadhesive properties to the hydrogel, facilitating the retention of hADSCs at cartilage defect sites following injection and enhancing therapeutic efficacy. Additionally, hADSCs within the 3D hydrogel matrix showed elevated gene expression linked to chondrogenic differentiation, and the hydrogel did not impede the expression of proteins associated with this process.

[0135] General Methods

[0136] Characterization of polymer chemical structures

[0137] Structural characterization and quantification of the degree of substitution were performed using proton nuclear magnetic resonance spectroscopy (’ H NMR, Bruker AV-500, Bruker BioSpin, Rheinstetten, Germany) based on established integration methods. To further characterize functional group incorporation, Fourier-transform infrared (FT-TR) spectroscopy was performed using a Nicolet 6700 instrument (Thermo Fisher Scientific). Polymer-KBr pellets were prepared by thoroughly mixing 1% (w / w) polymer with potassium bromide, followed by mechanical compression into translucent discs. In addition, ultra violet -visible (UV-Vis) spectroscopy (Agilent 8453, Agilent Technologies, Santa Clara, CA, USA) was used to verify successful catechol conjugation and to confirm the preservation of catechol hydroxyl functionality, as indicated by characteristic absorbance features and the absence of oxidation-associated peaks.

[0138] Morphological analysis of hydrogels

[0139] Hydrogels were fabricated with varying volume mixing ratios using DI water. To preserve their structural integrity, the prepared hydrogels were subjected to rapid cryogenic freezing using liquid nitrogen, followed by lyophilization to ensure complete dehydration. Morphological assessment was conducted via Scanning Electron Microscopy (SEM) at an accelerating voltage of 15 kV utilizing a Hitachi TM3000 Tabletop Microscope. For optimal imaging, specimens were affixed to double-sided carbon tape to enhance stability during electron beam interaction. Quantitative pore diameter measurements were obtained through digital image analysis using ImageJ software.

[0140] In vitro hydrogel degradation assessment

[0141] To evaluate the degradability of the hydrogel matrix, freshly synthesized hydrogels were subjected to freeze-drying to obtain their initial dry mass (Wo). The dehydrated specimens were subsequently immersed in a complete mesenchymal stem cell culture medium and maintained at 37 °C for predetermined intervals of 1, 3, 5, 7, 9, and 14 days. At each designated time point, the incubation medium was carefully aspirated, and the remaining hydrogel material was retrieved, lyophilized, and weighed (Wt). All experimental procedures were conducted in triplicate to ensure reproducibility. The percentage of residual hydrogel mass was determined using the following equation: (Wt / Wo) x 100.

[0142] In vitro cell culture

[0143] The complete cell culture medium was made by mixing 500 mL base medium, 25 mL FBS, 5 mL growth supplement, and 5 mL P / S solution. hADSCs were cultured in the complete medium in an incubator with 5% CO2 at 37 °C. The seeding density of hADSCs was 5000 cells / cm2in T-75 flask. The hADSCs were passaged every four days, and the cell culture medium was changed every two days. All the cells used were within eight passages.

[0144] Cell proliferation assay

[0145] Cell proliferation within hydrogels was evaluated using the Alamar Blue metabolic assay. hADSCs were encapsulated and hydrogel constructs were incubated in complete cell culture medium at 37 °C with 5% CO2. At designated time points, the medium was removed and replaced with 150 pL of fresh medium containing 10% (v / v) Alamar Blue reagent. After a 3 -hour incubation period, 100 pL of the resulting supernatant was transferred to a black 96-well fluorescence plate. Fluorescence intensity was measured using a microplate reader (SpectraMax M2, Molecular Devices, San Jose, CA, USA) with an excitation wavelength of 560 nm and emission detected at 590 nm. Signal intensity was used as an indicator of viable cell number and metabolic activity.

[0146] Statistical analysis

[0147] All graphs and bar charts were expressed as the mean ± standard deviation (SD) of more than five repeated experiments as described above. The mean values of the pore sizes of different hydrogels were compared by two-way analysis of variance (ANOVA), followed by Tukey’ s multiple comparisons test. Statistical comparisons of pore size across hydrogel groups, maximum adhesive strength, and relative gene expression levels were performed using one-way analysis of variance (ANOVA). The mean values of the size of the cell spheroid and the relative expression of RNA were compared by one- factor ANOVA and the significance of the difference was determined by Tukey’ s multiple comparisons test. Statistical analysis was conducted using Prism 9 software.

[0148] Materials and methods for the preparation of Suc-CS / Ald-HA hydrogels

[0149] Materials

[0150] Hyaluronic acid sodium salt from Streptococcus equi (about 1.5-1.8 x 106Da), succinic anhydride, sodium periodate, ethylene glycol, and tert-butyl carbazate were purchased from Sigma- Aldrich Inc. (St. Louis, MO, USA). Chitosan (medical grade, deacetylation degree: 95%) from Alaska snow crab was purchased from Matexcel (Bohemia, NY, USA). Calcein AM, propidium iodide, aggrecanmonoclonal antibody (BC-3), SOX9 recombinant rabbit monoclonal antibody (7H13L8), goat anti-mouse IgG, IgM (H+L) secondary antibody with Alexa Fluor™ 488, and donkey anti-rabbit IgG (H+L) highly cross-adsorbed secondary antibody with Alexa Fluor™ 555 were purchased from Thermo Fisher Scientific Inc. (Waltham, MA, USA). Human adipose-derived mesenchymal stem cells (hADSCs), MSC cell culture medium, fetal bovine serum (FBS), penicillin / streptomycin solution (P / S solution), and mesenchymal stem cell growth supplement were purchased from ScienCell Research Laboratories (Carlsbad, CA, USA). All chemicals and reagents were used as received.

[0151] Synthesis of TV-succinyl-chitosan (Suc-CS)

[0152] The water-soluble Suc-CS was synthesized through the chemical reaction between chitosan and succinic anhydride. Briefly, chitosan (0.25 g) was dissolved in 5% (v / v) acetic acid solution (20 mL), followed by the addition of methanol (80 mb). Succinic anhydride (0.75 g) was dissolved in acetone (10 mL) and added to the above solution. The reaction was performed at room temperature for 24 h under stirring. After the reaction, the pH of the solution was adjusted to 6 using 1 M NaOH, and the precipitates were collected through centrifugation (4000 ipm, 5 min) and re-dissolved in DI water. The solution was purified by 3-day dialysis against deionized water at room temperature using a Spectra / Por 3 RC dialysis membrane (3.5 K; Spectrum Laboratories, Inc., Rancho Dominguez, CA, USA). Finally, the purified solution was frozen using liquid nitrogen and dried using a lyophilizer (Labconco Co., Ltd., Kansas City, MO, USA). The purified polymer powder was stored at -20 °C. The chemical structure and substitution degree were analyzed and calculated using1H nuclear magnetic resonance spectroscopy (XH NMR, Bruker AV-500, Bruker BioSpin GmbH, Rheinstetten, Germany).

[0153] Synthesis of aldehyde hyaluronic acid (Ald-HA)

[0154] The hyaluronic acid was oxidized by sodium periodate to form Ald-HA. In brief, hyaluronic acid (2 g) was dissolved in DI water (200 mL) and the mixture was stirred overnight in a 500 mL flask. Sodium periodate solution in DI water (0.5 M, 10 mL) was then added to the solution, and the reaction was performed at room temperature for 2 h, protected from light. Ethylene glycol (2 mL) was then added to the reactant and stirred at room temperature for 1 h to stop the reaction. The solution was purified by 3-day dialysis against deionized water at room temperature using a Spectra / Por 3 RC dialysis membrane. Finally, the purified solution was frozen using liquid nitrogen and dried using a lyophilizer.The purified polymer powder was stored at -20 °C. The chemical structure was analyzed using ’ H NMR. The aldehydes content in Ald-HA was quantified using tert-butyl carbazate. The peak corresponding to the tertbutyl substituent ((CfUpCOCONHNH-, 5 = 1.38 ppm) was compared with the peak of HA acetamide protons at 1.9 ppm.

[0155] Preparation of Suc-CS / Ald-HA hydrogels

[0156] The Suc-CS and Ald-HA were dissolved separately into DPBS at different concentrations (10, 20, 30, 40, and 50 mg / mL). The Suc-CS / Ald-HA hydrogels were prepared by mixing the two polymer solutions of the same concentration (1 / 1, v / v) at room temperature using a double barrel syringe (Duploject syringe, Baxter Healthcare, Deerfield, IL, USA) with a 21 G needle (Becton & Dickinson, Franklin Lakes, NJ). For morphological observation, DPBS was replaced by distilled water to prepare the hydrogel.

[0157] Self-healing capability of Suc-CS / Ald-HA hydrogels

[0158] 700 pL of freshly prepared hydrogels was cut into two pieces and one of them (350 pL) was soaked in 500 pL Trypan Blue for 5 min. Then, two pieces of hydrogel were brough together at room temperature for 3 min. The resulting hydrogel piece was picked up using a tweezer.

[0159] Rheological analysis of Suc-CS / Ald-HA hydrogels

[0160] The rheological properties of crosslinked Suc-CS / Ald-HA hydrogels of different concentrations (10, 20, 30, 40, and 50 mg / mL) were evaluated using a rheometer (Anton Paar MCR 92, Anton Paar GmbH, Graz, Austria). For stress sweeps, the storage (G') and loss modulus (G”) were recorded under increasing shear stress from 0.1-10000 Pa at temperature of 37 °C and frequency of 1 Hz. For time sweep, the storage (G’) and loss (G”) modulus were recorded over time for Suc-CS / Ald-HA hydrogels of different concentrations (10, 20, 30, 40, and 50 mg / mL) by rheological analysis at temperature of 37 °C, shear strain of 1%, and frequency of 1 Hz.

[0161] Suc-CS / Ald-HA hydrogel gelation times

[0162] The inverting tube method was used to determine the gelation times of hydrogels. The polymer solutions of Suc-CS and Ald-HA were incubated at 37 °C before mixing. The Suc-CS and Ald-HA mixture was incubated at 37 °C post mixing, and its flowability was assessed. The gelation time was defined as the point at which the hydrogel solution ceased to flow. The evaluation was conducted in triplicate.

[0163] Suc-CS / Ald-HA hydrogel morphology

[0164] Suc-CS and Ald-HA polymers were separately dissolved in DI water at different polymer concentrations (10, 20, 30, 40, and 50 mg / mL). Suc-CS / Ald-HA hydrogels were prepared upon mixing two polymer solutions of the same concentration (1 / 1, v / v) at room temperature. 200 pL of the hydrogel was soaked into 300 pL DI water and incubated at 37 °C for 1, 3, and 7 days. After the incubation, the solution was removed using a pipette and the remaining hydrogel were completely soaked into liquid nitrogen for 30 min and then loaded on the lyophilizer for 2 days. The freeze-dried samples were observed using a SEM with an acceleration voltage of 15kV (Hitachi TM3000 Tabletop Microscope, Hitachi High-Technologies Corporation, Tokyo, Japan). The pore size was determined by analyzing SEM images using ImageJ. First, the SEM image was imported into ImageJ, and a scale was established based on known dimensions. Second, appropriate tools such as the line tool were utilized to identify individual pores and quantify their diameters. Third, the average pore size was derived from multiple measurements.

[0165] Live / dead staining: Suc-CS / Ald-HA hydrogels

[0166] Calcein-AM and propidium iodide were used to stain live and dead cells, respectively. Calcein-AM was dissolved in dimethyl sulfoxide (DMSO) to obtain a stock concentration of 1 mM, and propidium iodide was dissolved in DI water to obtain a stock concentration of 1.5 mM. 96-well plates were used for Eive / Dead staining test. First, hADSCs were suspended in the Ald-HA solution. Then, 75 pL of the Suc-CS solution was mixed with 75 pL of hADSCs-Ald-HA solution in each well of a 96-well plate using a pipette. The total volume of hydrogel in each well was 150 pL, and the total number of hADSCs in each well was 10,000 cells. The concentrations of Suc-CS and Ald-HA were: 10 mg / mL, 20 mg / mL, 30 mg / mL, 40 mg / mL, and 50 mg / mL in a complete cell culture medium. The experiment for each concentration was repeated three times. The hydrogel was incubated at 37 °C for 30 min, then 100 pL complete medium was added on top of the hydrogel. hADSCs in the hydrogel were cultured for 1, 3, 5, 7, and 14 days. The cell culture medium was changed every day, and DPBS was added to the edge of plate to reduce the evaporation. The live / dead staining solution was made by mixing 4 pL of Calcein-AM (2 mM, DMSO), 4 pL PI (2 mM, DI), and 1 mL DPBS. Before adding the staining solution to the hydrogel, the cell culture medium was removed and the hydrogel was gently rinsed with DPBS twice. After removing the DPBS, 100 pL staining solution was added to each well and the solution was incubated for 15 min at room temperature. The dyeing solutionwas then replaced with an equal volume of DPBS. A fluorescence microscope (Olympus 1X81, Olympus Life Science) was used to observe and take images of cells.

[0167] cDNA synthesis and quantitative reverse transcriptase polymerase chain reaction (qRT-PCR)

[0168] To quantify relative gene expression, qRT-PCR was performed on complementary deoxyribonucleic acid (cDNA) collected from cells cultured in both hydrogels and 2D surfaces after 7 days and 14 days of growth. 24-well plates were used for PCR test. First, hADSCs were suspended in the Ald-HA solution. Then, 350 pL of hADSCs-Ald-HA solution was mixed with 350 pL of Suc-CS solution in each well of a 24-well plate using a pipette. The total volume of hydrogel in each well was 700 pL, and the total number of hADSCs in each well was 140,000 cells. The concentrations of Suc- CS and Ald-HA were: 20 mg / mL and 30 mg / mL, and the experiment for each concentration was repeated three times. Cells were cultured in the chondrogenic differentiation medium for 7 days and 14 days. The medium was changed every day. The Qiagen RNeasy Kit (Qiagen Inc., Valencia, CA, USA) was used to extract ribonucleic acid (RNA) from cells and the iScript cDNA synthesis kit was used to make cDNA according to the manufacturer’s instruction. In a Bio-Rad CFX96 real-time PCR detection system (BioRad Laboratories, Inc. Hercules, CA, USA), the amplification with a primer for each of the transcripts was evaluated using SYBR Green PCR Mastermix. The reference gene was glyceraldehyde 3-phosphate dehydrogenase (GAPDH). The cycle number at threshold fluorescence intensity was used for quantification. All samples were thermocycled in a 10 pL solution comprising 5 pL SYBR Mastermix, 300 nm primers, Integrated DNA Technologies, Coralville, IA), and cDNA at 0.04-1 ng / pL concentration adjusted for each condition. The thermocycle was 95 °C for 2 min, followed by 40 cycles of denaturation at 95 °C for 15 s, annealing at 58 °C for 30 s, and extension at 72 °C for 30 s. The CFX Manager program was used to examine all RT-qPCR results (Bio-Rad Laboratories, Inc.).

[0169] Immunofluorescence and immunocytochemistry: Suc-CS / Ald-HA hydrogels

[0170] The samples were prepared using the identical procedure as for the Live / Dead staining method. Specifically, 35 pL of Suc-CS solution was mixed with 35 pL of hADSCs-Ald-HA solution in each well of a 96-well plate using a pipette. The total volume of hydrogel in each well was 70 pL, and the total number of hADSCs in each well was 14,000 cells. The concentrations of Suc-CS and Ald-HA were 20 mg / mL and 30mg / mL, and the experiment for each concentration was repeated three times. Subsequently, the cells were cultivated in a medium designed for chondrogenic differentiation for 14 days, with daily medium replacement. The cells within the hydrogel were fixed with a 4% paraformaldehyde solution for 15 min at room temperature. Following fixation, the cells were pemieabilized with 1% (v / v) Triton X-100 in DPBS and blocked with 15% donkey serum albumin and 0.3% Triton X-100 in DPBS for 1 h at room temperature. The cells were then incubated with primary antibodies ACAN and SOX9 overnight at 4 °C, followed by incubation with the corresponding secondary antibodies for 1 h at room temperature in the dark. The nuclei were counterstained with 1 pM DAPI for 15 min. Between each of these steps, thorough washing with cold PBS was earned out at least three times to remove any unbound antibodies or other reagents. All the immunostained samples were imaged using a Nikon TE300 inverted microscope (Nikon Corporation, Tokyo, Japan).

[0171] Materials and methods for preparation of Ald-HA / Dopa-HBCS hydrogels

[0172] Materials

[0173] The primary reagents used in this study included chitosan with a medium molecular weight (190-310 kDa, 75-85% deacetylated) and hyaluronic acid (HA) sodium salt derived from Streptococcus equi(molecular weight about 1.5-1.8 x 106Da), both obtained from Sigma-Aldrich (St. Louis, MO, USA). Additional synthetic and analytical- grade chemicals sourced from Sigma- Aldrich included potassium hydroxide (KOH), 1,2- epoxybutane, sodium periodate, ethylene glycol, deuterium oxide, N-hydroxysuccinimide (NHS), and l-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC HC1). Calcium chlorides were supplied by Avantor (Radnor, PA, USA). Thermo Fisher Scientific (Waltham, MA, USA) provided urea, 3,4-dihydroxyphenylacetic acid, live / dead staining dyes (calcein AM and propidium iodide), and all antibodies, including aggrecan monoclonal antibody (clone BC-3), SOX9 rabbit monoclonal antibody (clone 7H13L8), Alexa Fluo™ 488-labeled goat anti-mouse IgG / IgM (H+L), and Alexa Fluor™ 555- labeled donkey anti-rabbit IgG (H+L, highly cross-adsorbed). Anhydrous ethanol was purchased from Decon Laboratories (King of Prussia, PA, USA). All materials were used without further modification.

[0174] Synthesis of hydroxybutyl chitosan (HBCS)

[0175] HBCS was synthesized through a version of a published protocol (Bi et al., Temperature sensitive self-assembling hydroxybutyl chitosan nanoparticles with cationic enhancement effect for multi-functional applications. Carbohydrate Polymers2021, 254, 117199). In brief, 300 mg of chitosan was dispersed in an aqueous alkaline medium prepared by dissolving 3.2 g of KOH and 1.6 g of urea in 14.6 g of deionized (DI) water. The suspension was stirred at ambient temperature for 30 minutes, followed by two freeze-thaw cycles at -20 °C for 3 hours each to enhance dissolution. After filtration to remove undissolved particulates, a clear and uniform chitosan solution was obtained. To introduce hydroxybutyl groups, the resulting chitosan solution was combined with 1,2- epoxybutane at a volume ratio of 1 :4 and stirred for 48 hours at room temperature. The reaction mixture was then adjusted to neutral pH using hydrochloric acid, followed by purification through dialysis (Spectra / Por 3 RC, MWCO 3.5 kDa; Spectrum Laboratories, Rancho Dominguez, C A) against DI water for two days at room temperature. The dialyzed product was frozen in liquid nitrogen and lyophilized (Labconco, Kansas City, MO) to yield dry HBCS powder. The final polymer powder was stored at -20 °C until use.

[0176] Synthesis of catechol-modified HBCS (Dopa-HBCS)

[0177] Dopa-HBCS was prepared via carbodiimide-mediated coupling, following a version of an established protocol (Shou et al., Thermoresponsive chitosan / DOPA-based hydrogel as an injectable therapy approach for tissue-adhesion and hemostasis. ACS Biomaterials Science & Engineering 2020, 6 (6), 3619-3629). A total of 190 mg of HBCS was dissolved in 10 mL of MES buffer (pH 4.48), and the solution was subsequently degassed under vacuum to eliminate dissolved oxygen. To prevent oxidative degradation of catechol groups, all subsequent steps were conducted under a nitrogen atmosphere. 3,4-Dihydroxyphenylacetic acid was introduced to the deoxygenated HBCS solution, followed by the dropwise addition of a freshly prepared 50% aqueous ethanol solution (v / v) containing equimolar concentrations of EDC-HC1 and NHS. The reaction mixture was stirred at room temperature for 24 hours under continuous nitrogen flow. Upon completion, unreacted small molecules and byproducts were removed by dialysis against deionized water (pH 5.0 ± 0.05) for two days using a Spectra / Por 3 RC membrane (MWCO 3.5 kDa) at ambient temperature. The resulting catechol-modified HBCS was frozen in liquid nitrogen and lyophilized to obtain dry Dopa-HBCS.

[0178] Synthesis of aldehyde-functionalized hyaluronic acid (Ald-HA)

[0179] Ald-HA was synthesized via periodate oxidation of the vicinal diols on the HA backbone. Specifically, 1.0 g of HA was dissolved in 100 mL of DI water and stirred until fully solubilized. Then 5 mL aqueous sodium periodate (0.5 M in DI) was slowly added to the HA solution under dark conditions to prevent light-induceddecomposition. The reaction proceeded at room temperature for 2 hours with continuous stirring. To quench any unreacted periodate, 1 mL of ethylene glycol was added, and the mixture was stirred for an additional hour. The reaction product was subsequently purified by dialysis against deionized water for three days at ambient temperature using a Spectra / Por 3 RC dialysis membrane (MWCO 3.5 kDa; Spectrum Laboratories, USA). The dialyzed solution was frozen in liquid nitrogen and lyophilized to yield dry Ald-HA. The final Ald-HA product was stored as a dry powder at -20 °C until use.

[0180] Preparation of Ald-HA / Dopa-HBCS hydrogels

[0181] Ald-HA and Dopa-HBCS were dissolved separately into DI or DPBS at different concentrations of 5, 10, 20, or 30 mg / mL. The Dopa-HBCS solutions were supplemented with 100 mM CaCb. Hydrogels were formed by mixing equal-concentration solutions of Ald-HA and Dopa-HBCS at predefined volume ratios (3:7, 5:5, and 7:3) using a dual-barrel syringe (Duploject, Baxter International Inc., IL, USA) fitted with a 21-gauge needle (Becton & Dickinson, Franklin Lakes, NJ, USA).

[0182] Rheological characterization of Ald-HA / Dopa-HBCS hydrogels

[0183] The viscoelastic behavior of Ald-HA / HBCS and Ald-HA / Dopa-HBCS hydrogels was evaluated using a rotational rheometer (Anton Paar MCR 92, Virginia, USA) equipped with a 25 mm parallel-plate (PP25) geometry. Temperature control was maintained via an integrated air-based heating and cooling system, and all measurements were conducted at 37 °C unless otherwise specified. Hydrogels were freshly prepared following the procedure described above and immediately transferred to the rheometer platform. To assess mechanical performance, an amplitude sweep test was first conducted by applying increasing shear stress from 0.1 to 10000 Pa at a constant frequency of 1 Hz, allowing determination of the storage modulus (G’) and loss modulus (G”) across hydrogels of varying polymer concentrations. Subsequently, a time sweep test was performed to monitor gelation kinetics and initial structural stability post-injection. Measurements began immediately after mixing, and G’ and G” were recorded over time at a fixed strain of 1% and frequency of 1 Hz, simulating physiological conditions and enabling evaluation of early-stage mechanical integrity and potential for cell retention postinjection.

[0184] Self-healing capability of Ald-HA / Dopa-HBCS hydrogels

[0185] 200 pL of freshly prepared Ald-HA / Dopa-HBCS hydrogel was cut into two pieces and one of them (100 pL) was soaked in 500 pL Trypan Blue for 5 min. Then,two pieces of hydrogel were brough together at room temperature for 3 min. Finally, half of the formed hydrogel was adhered to an inclined glass slide to test the self-healing capability. The self-healing process was further assessed using time-sweep rheological measurements at 37°C, 1% shear strain, and 1 Hz frequency. The recovery ratio of the storage modulus (G’) was calculated as (G’t / G’o) x 100, where G’t represents G’ at different time points after the hydrogel fragments were rejoined, and G’o represents G’ before cutting.

[0186] In vitro cell experiments: Ald-HA / Dopa-HBCS hydrogels

[0187] Human adipose-derived mesenchymal stem cells (hADSCs) and associated reagents, including basal mesenchymal stem cells culture medium, fetal bovine serum (FBS), penicillin-streptomycin (P / S), and mesenchymal stem cell growth supplement, were obtained from Sciencell (Carlsbad, CA, USA). Complete growth medium was prepared by supplementing 500 mL of base medium with 25 mL FBS, 5 mL growth supplement, and 5 mL P / S solution. Cells were maintained at 37 °C in a humidified incubator with 5% CO? and subcultured every four days. Medium was refreshed every other day. For hydrogel preparation and cell encapsulation, Ald-HA and Dopa-HBCS polymers were individually dissolved in complete culture medium containing 100 mM CaCh. All polymer solutions were sterilized by filtration through 0.2 pm syringe filters and adjusted to a final concentration of lO mg / mL. Hydrogels were formulated at two different volume mixing ratios (5:5 and 3:7) for both the Ald-HA / HBCS and Ald- HA / Dopa-HBCS systems. Specifically, the A5H5 hydrogel was formed by combining equal volumes of lO mg / mL Ald-HA and lO mg / mL HBCS, while A3H7 was prepared using a 3:7 volume ratio. Analogously, A5D5 and A3D7 referred to the corresponding mixtures incorporating Dopa-HBCS in place of unmodified HBCS.

[0188] Live / Dead cell viability assay: Ald-HA / Dopa-HBCS hydrogels

[0189] Cell viability within hydrogels was assessed using a Live / Dead staining protocol based on calcein-AM and propidium iodide (PI). Stock solutions were prepared by dissolving calcein-AM in DMSO (1 mM) and PI in deionized water (1 .5 mM). Polymer solutions of Ald-HA and Dopa-HBCS (each at lO mg / mL) were filtered through 0.2 pm syringe filters and prepared in complete mesenchymal stem cells culture medium containing 100 mM CaCh. For hydrogel formation, Dopa-HBCS solution was first dispensed into 96-well plates. A defined number of hADSCs was suspended in Ald-HA solution and subsequently mixed with the Dopa-HBCS solution at either 5:5 or 3:7 volumeratios to yield a final volume of 70 pL per well, containing 14,000 cells. Following gelation at 37 °C for 30 minutes, 150 pL of fresh culture medium was added to each well and replenished every two days. At the indicated time points, cell-laden hydrogels were stained by adding 100 pL of Live / Dead dye solution (2 pL calcein-AM + 2 pL PI in 1 mL DPBS) and incubated for 15 minutes at 37 °C. The staining solution was then replaced with DPBS prior to imaging. Fluorescence images were captured using an Olympus 1X81 inverted fluorescence microscope (Olympus Life Science). In addition, cell viability following syringe injection was also evaluated using the same staining protocol to assess potential mechanical damage during delivery.

[0190] Tissue adhesion test

[0191] The adhesive strength of the hydrogels to biological tissue was evaluated using a lap shear configuration with fresh porcine skin as the substrate. Skin samples were thoroughly rinsed with DPBS and 75% ethanol, then cut into uniform rectangular strips measuring 15 x 40 x 2 mm. Prior to testing, polymer solutions of Ald-HA and Dopa-HBCS (lO mg / mL each) were dissolved in deionized water containing 100 mM CaCh and sterilized by filtration. A total of 200 pL of mixed polymer solution was injected between two overlapping skin sections (15 x 15 mm contact area) using a dual-barrel syringe. The assembled constructs were incubated at 37 °C overnight to allow complete gelation and tissue bonding. Shear adhesion strength was quantified using a mechanical testing system (AGS-X, Shimadzu Corporation, Kyoto, lapan) by pulling the skin pieces apart at a constant rate of 1.3 mm / min until failure occurred. The maximum load at detachment was recorded for comparison between hydrogel formulations.

[0192] Gene expression characterization of hADSC chondro genesis within Ald-HA / Dopa-HBCS hydrogels

[0193] Gene expression analysis was performed to assess the chondrogenic differentiation of hADSCs encapsulated within hydrogels. For each sample, 140,000 cells were suspended in Ald-HA solution and combined with an equal volume of Dopa-HBCS solution (350 pL each), yielding a total hydrogel volume of 700 pL per well in a 24-well plate. Constructs were cultured in chondrogenic differentiation medium for 7 or 14 days, with media changes every two days. At each time point, total RNA was extracted using the RNeasy Mini Kit (Qiagen, Valencia, CA, USA), and complementary DNA (cDNA) was synthesized using the iScript cDNA synthesis kit according to the manufacturer’s instructions. Quantitative PCR was carried out on a CFX96 real-time PCR detectionsystem (Bio-Rad, Hercules, CA, USA) using SYBR Green Master Mix and gene-specific primers (Integrated DNA Technologies, Coralville, 1A, USA). Glyceraldehyde 3- phosphate dehydrogenase (GAPDH) was used as the reference gene. The amplification protocol consisted of an initial denaturation at 95 °C for 2 min, followed by 40 cycles of 95 °C for 15 s, 58 °C for 30 s, and 72 °C for 30 s. Relative gene expression levels were calculated based on threshold cycle (Ct) values using the Bio-Rad CFX Manager software.

[0194] Immunofluorescence and immunocytochemistry: Ald-HA / Dopa-HBCS hydrogels

[0195] Immunofluorescence staining was conducted to visualize protein-level expression of chondrogenic markers in hADSCs encapsulated within hydrogels. In each well of a 96-well plate, 35 pL of Dopa-HBCS solution was combined with 35 pL of hADSC-containing Ald-HA solution, yielding 70 pL of hydrogel containing 14,000 cells. Both polymer components were prepared at 10 mg / mL and filtered through 0.2 pm syringe filters prior to use. Cell-laden hydrogels were cultured in chondrogenic differentiation medium for 14 days, with media replenished every two days. Following the culture period, samples were fixed in 4% paraformaldehyde for 15 minutes at room temperature, then permeabilized using 1% Triton X-100 in DPBS. Nonspecific binding sites were blocked using a solution of 15% donkey serum and 0.3% Triton X-100 in DPBS for 1 hour. Primary antibodies against ACAN and SOX9 were applied and incubated overnight at 4 °C. After thorough washing with cold PBS, samples were incubated with fluorescently labeled secondary antibodies for 1 hour in the dark at room temperature. After incubation with 1 pM DAPI for 15 minutes to label cell nuclei, the stained hydrogel constructs were visualized using an inverted fluorescence microscope (Nikon TE300, Nikon Corporation, Tokyo, Japan).

[0196] As used herein, the term “about” refers to ± 5% of the specified value.

[0197] While illustrative embodiments have been illustrated and described, it will be appreciated that various changes can be made therein without departing from the spirit and scope of the invention.

Claims

CLAIMSThe embodiments of the invention in which an exclusive property or privilege is claimed are defined as follows:

1. A method for delivering stem cells for cartilage regeneration, comprising substantially simultaneously administering a first composition comprising a chitosanhydrogel precursor and a second composition comprising a hyaluronic acid-hydrogel precursor to a site in need of cartilage regeneration to provide an in situ stem cellcontaining chitosan / hyaluronic acid hydrogel at the site, wherein the first composition, the second composition, or both the first and the second compositions further comprise stem cells.

2. The method of Claim 1, wherein the first composition comprises stem cells.

3. The method of Claim 1, wherein the second composition comprises stem cells.

4. The method of Claim 1, wherein the stem cells are human adipose-derived stem cells.

5. The method of Claim 1, wherein the chitosan-hydrogel precursor is a chitosan polymer comprising a tissue adhesive component.

6. The method of Claim 1, wherein the hyaluronic acid-hydrogel precursor is a hyaluronic acid polymer having aldehyde groups.

7. The method of Claims 5 or 6, wherein the stem cell-containing chitosan / hyaluronic acid hydrogel is formed by Schiff base reaction between the amino groups of the chitosan polymer and the aldehyde groups of the hyaluronic acid polymer.

8. The method of any one of Claims 1-7, wherein the hydrogel exhibits tissueadhesive properties.

9. The method of Claim 8, wherein the tissue-adhesive properties are tunable by adjusting the concentration of the chitosan-hydrogel precursor, the hyaluronic acidhydrogel precursor, or both.

10. The method of Claim 9, wherein the chitosan-hydrogel precursor composition and the hyaluronic acid-hydrogel precursor composition have polymer concentrations between 5 mg / mL and 30 mg / mL.

11. The method of any one of Claims 1-10, wherein the hydrogel exhibits self- healing behavior following mechanical disruption.

12. The method of Claim 11, wherein the self-healing behavior arises from reversible crosslinking interactions between the chitosan-based and hyaluronic acid-based precursors.

13. The method of Claim 11, wherein the hydrogel recovers at least 98% of its storage modulus within 3 minutes after mechanical disruption.

14. The method of any one of Claims 1-13, wherein the hydrogel maintains stem cell viability following injection and in situ gelation.

15. The method of any one of Claims 1-13, wherein the hydrogel supports chondrogenic differentiation of the stem cells.

16. The method of any one of Claim 1-13, wherein the first and second compositions are administered by injection.

17. The method of any one of Claims 1-16, wherein the site is a knee.

18. A chitosan / hyaluronic acid hydrogel formed in situ in vivo, comprising one or more chitosan polymers covalently coupled to one or more hyaluronic acid polymers by Schiff base reactions, wherein the hydrogel comprises stem cells.

19. The hydrogel of Claim 18, wherein the stem cells are human adipose- derived stem cells.

20. The hydrogel of Claim 18, wherein the one or more chitosan polymers comprise a tissue adhesive component.

21. The hydrogel of Claim 18, wherein the one or more hyaluronic acid polymers are hyaluronic acid polymers having aldehyde groups.

22. The hydrogel of any one of Claims 18-21, wherein the hydrogel exhibits tissue-adhesive properties.

23. The hydrogel of Claim 22, wherein the tissue-adhesive properties are tunable by adjusting the concentration of the one or more chitosan polymer, the one or more hyaluronic acid polymers, or both.

24. The hydrogel of any one of Claims 18-21, wherein the hydrogel exhibits self-healing behavior following mechanical disruption.

25. The hydrogel of Claim 24, wherein the self-healing behavior arises from reversible crosslinking interactions between the chitosan-based and hyaluronic acid-based precursors.

26. The hydrogel of Claim 24, wherein the hydrogel recovers at least 98% of its storage modulus within 3 minutes after mechanical disruption.

27. The hydrogel of any one of Claims 18-26, wherein the hydrogel maintains stem cell viability following injection and in situ gelation.

28. The hydrogel of any one of Claims 18-26, wherein the hydrogel supports chondrogenic differentiation of the stem cells.

29. The hydrogel of any one of Claims 18-26, wherein the hydrogel comprises pores having an average pore size between about 15 pm and about 35 pm.

30. A kit for delivering stem cells for cartilage regeneration comprising:(a) a first composition comprising a chitosan-hydrogel precursor;(b) a second composition comprising a hyaluronic acid-hydrogel precursor; and(c) stem cells.

Images