Acellular matrix double-network hydrogel as well as preparation method and application thereof

Abstract

The invention discloses acellular matrix double-network hydrogel as well as a preparation method and application thereof. Specifically, the invention provides the acellular matrix double-network hydrogel, the acellular matrix double-network hydrogel is provided with a first network formed by thermal gelation of an acellular matrix and a second network formed by (methyl) acryloylated gelatin, and the first network and the second network are interspersed to form the double-network hydrogel. The hydrogel has excellent biocompatibility, controllable physical properties and active biological signal regulation and control capability, can be used as a platform material, and has excellent application prospects in the fields of tissue repair and regenerative medicine, especially in the aspects of nerve regeneration and stem cell proliferation and induced differentiation.

Description

Technical Field

[0001] This invention belongs to the field of biomedical materials and tissue engineering technology, specifically relating to a decellularized matrix dual-network hydrogel, its preparation method, and its application. Background Technology

[0002] Microcarriers, due to their high specific surface area and excellent cell suspension culture capabilities, are widely used in tissue engineering, drug screening, and regenerative medicine. Currently, materials used to construct microcarriers are mainly divided into synthetic polymers (such as polylactic acid PLA and polycaprolactone PCL) and natural polymers (such as gelatin, alginate, and chitosan). Synthetic polymers typically possess good mechanical strength but often lack cell recognition sites; while single-component natural polymers, although exhibiting good biocompatibility, struggle to fully mimic specific tissues, especially the nervous system. The complex extracellular microenvironment (ECM) also presents limitations in inducing specific differentiation of stem cells.

[0003] Decellularized matrix (dECM) materials can provide a tissue-specific biomimetic microenvironment by removing the cellular components of native tissue while retaining collagen, glycosaminoglycans, and various signaling factors. However, dECM materials, especially those derived from spinal cord or brain tissue, typically exhibit soft texture and low elastic modulus.

[0004] However, the mechanical properties of existing single dECM microcarriers are often fixed and unadjustable, making it difficult to precisely match the differentiated mechanical support requirements of different repair stages. Therefore, there is a need in this field for a decellularized matrix material that can both retain the rich biological activity of dECM and adapt to the needs of different tissues. Summary of the Invention

[0005] This invention provides a heparin-loaded decellularized matrix dual-network hydrogel microcarrier and its preparation method.

[0006] In a first aspect, the present invention provides a method for preparing a decellularized matrix dual-network hydrogel, comprising the following steps: S1. Preparation of aqueous precursor solution: Dissolve decellularized matrix (dECM) in water under acidic conditions, then mix with an aqueous solution of (meth)acrylamide gelatin and photoinitiator, and adjust the pH to neutral at a low temperature of 0-10℃ to obtain an aqueous precursor solution; S2. Primary thermogelation: The aqueous precursor solution is heated to 35-40°C to induce primary thermogelation of the decellularized matrix, forming a mixture containing a first network gel; and S3. Photocrosslinking and curing: Applying light of a predetermined wavelength to the mixture containing the first network gel initiates a polymerization reaction of (meth)acrylamide gelatin to form a second network, thereby obtaining an interpenetrating double network hydrogel.

[0007] In another preferred embodiment, the method further includes: S4. Functional modification: covalently grafting heparin molecules onto the amino groups on the surface of the dual-network hydrogel via an amidation reaction.

[0008] In another preferred embodiment, in step S1, the (meth)acrylamide gelatin is selected from methacrylamide gelatin and acrylamide gelatin, preferably methacrylamide gelatin.

[0009] In another preferred embodiment, the concentration of decellularized matrix in the aqueous precursor solution is 5-50 mg / mL (0.5-5.0 w / v%), preferably 7.5-20 mg / mL, more preferably 10-15 mg / mL, such as 8 mg / mL, 9 mg / mL, 9.5 mg / mL, 10 mg / mL, 10.5 mg / mL, 11 mg / mL, 12 mg / mL, 13 mg / mL, 14 mg / mL or 15 mg / mL.

[0010] In another preferred embodiment, the concentration of (meth)acrylamide gelatin in the aqueous precursor solution is 1-20 w / v, preferably 2-10 w / v, more preferably 3-6 w / v, such as 3 w / v%, 3.5 w / v%, 4 w / v%, 4.5 w / v%, 5 w / v%, 6 w / v%, 7 w / v%, 8 w / v%, 9 w / v%, or 10 w / v.

[0011] In another preferred embodiment, the photoinitiator is selected from one or more of lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP), lithium 2,4,6-trimethylbenzoyl-diphenylphosphine oxide (Li-TPO), camphorquinone (CQ), riboflavin, or the group consisting of lithium phenyl-2,4,6-trimethylbenzoyl-diphenylphosphine oxide (Li-TPO), camphorquinone (CQ), and riboflavin.

[0012] In another preferred embodiment, the (meth)acrylamide gelatin has an average molecular weight of 50 kDa to 100 kDa.

[0013] In another preferred embodiment, the amount of photoinitiator is 0.1-3% w / v of the volume of the (meth)acrylamide gelatin solution, preferably 0.2-2% w / v, more preferably 0.4-1% w / v, such as 0.5% w / v or 0.8% w / v.

[0014] In another preferred embodiment, the (meth)acrylamide gelatin (GelMA) is prepared by reacting type A gelatin with methacrylic anhydride or acrylic anhydride.

[0015] In another preferred embodiment, the type A gelatin is derived from pigskin, bovine bones, and cowhide.

[0016] In another preferred embodiment, the (meth)acrylamide gelatin has a Bloom value of 300-350, preferably 310-330, and more preferably 320-325.

[0017] In another preferred embodiment, the (meth)acrylamide gelatin is one or more with a degree of substitution selected from 5-95%, such as one or more with a degree of substitution of 10%, 20%, 30%, 40%, 50%, 60%, 70%, 80%, 90%, and 95%.

[0018] In another preferred embodiment, the (meth)acrylamide gelatin is a mixture of (meth)acrylamide gelatin with low and high degrees of substitution. This is advantageous for preparing dual-network hydrogels with low stress relaxation times.

[0019] In another preferred embodiment, the (meth)acrylamide gelatin is a mixture of (meth)acrylamide gelatin with a degree of substitution of 10%-20% and 85%-95%; preferably, the mass ratio of the (meth)acrylamide gelatin with a degree of substitution of 10%-20% and 85%-95% in the mixture is 1:(0.5-2), preferably 1:(0.8-1.5), such as 1:0.6, 1:0.7, 1:0.8, 1:0.9, 1:1, 1:1.1, 1:1.2, 1:1.3, 1:1.4, 1:1.5, 1:1.6, 1:1.7, 1:1.8, 1:1.9 or 1:2.

[0020] In another preferred embodiment, in step S1, the acidic condition refers to a pH of 1-4, preferably pH 2-3.

[0021] In another preferred embodiment, in step S1, the acidic condition refers to a 0.008-0.15M HCl solution, preferably a 0.01-0.12M HCl solution.

[0022] In another preferred embodiment, in step S1, the low temperature is -4 to 10°C, preferably 0 to 4°C; In another preferred embodiment, in step S1, adjusting the pH to neutral means a pH value of 6.8-7.4, preferably adjusted with NaOH.

[0023] In another preferred embodiment, in step S1, the solvent in the aqueous mixed solution of (meth)acrylamide gelatin and photoinitiator is water or an aqueous physiological buffer, such as phosphate buffer.

[0024] In another preferred embodiment, step S2 includes the following steps: using the aqueous precursor solution as the dispersed phase and an oily liquid containing a surfactant as the continuous phase, droplet generation is performed in a microfluidic chip, and primary thermal gelation of the decellularized matrix is ​​induced at the rear end of the chip by heating to form a mixture microsphere containing a first network gel.

[0025] In another preferred embodiment, the method includes the steps of: S1. Preparation of aqueous precursor solution: Dissolve the decellularized matrix in water under acidic conditions, then mix it with an aqueous solution of (meth)acrylamide gelatin and photoinitiator, adjust the pH to neutral at a low temperature of 0-10℃ to obtain an aqueous precursor solution, and store it at a low temperature of 0-10℃. S2. Primary thermogelation: Using the aqueous precursor solution as the dispersed phase and an oily liquid containing surfactants as the continuous phase, droplet generation is performed in a microfluidic chip. Primary thermogelation of the decellularized matrix is ​​induced at the rear end of the chip by heating to 35-40℃, forming microspheres containing a first network gel. S3. Photocrosslinking and curing: Apply light of a predetermined wavelength to the mixture microspheres containing the first network gel to initiate the polymerization reaction of (meth)acrylamide gelatin to form a second network, thereby obtaining interpenetrating double-network hydrogel microspheres.

[0026] In another preferred embodiment, the oily liquid is selected from one or more of liquid paraffin, mineral oil, vegetable oil, and fluorinated oil.

[0027] In another preferred embodiment, the surfactant is selected from one or more of Span, Tween, and fluorinated surfactants, preferably Span 80.

[0028] In another preferred embodiment, the amount of surfactant added in the continuous phase is 0.1-10.0 w / v%, preferably 2-8 w / v%, more preferably 4-6 w / v.

[0029] In another preferred embodiment, in step S2, the heating temperature is 37±1℃.

[0030] In another preferred embodiment, in step S2, during the droplet generation process, the flow rate ratio of the dispersed phase to the continuous phase is 1:(1-20), preferably 1:(2-10), and more preferably 1:(5-8).

[0031] In another preferred embodiment, the diameter of the microspheres is 100-300 μm, preferably 150-250 μm.

[0032] In another preferred embodiment, in step S3, the predetermined wavelength illumination is blue light of 365-480 nm, and the illumination intensity is 10-100 mW / cm².2 The illumination time is 10-120 seconds; preferably, it is blue light at 380-450 nm with an intensity of 20-50 mW / cm². 2 The illumination time is 30-80 seconds.

[0033] In another preferred embodiment, in step S4, the amidation reaction is performed by activating the carboxyl group in the heparin molecule with EDC / NHS and then reacting it with amino groups on the surface of the dual-network hydrogel (including amino groups on the decellularized matrix and (meth)acrylamide gelatin).

[0034] In another preferred embodiment, before S4, the following steps may be included: adding a demulsifier and washing to remove the oil phase by centrifugation or sieving, while protecting the spherical morphology of the microspheres.

[0035] In another preferred embodiment, the heparin grafting density in the dual-network hydrogel is 0.1-5 μg / mg, preferably 0.5-4 μg / mg, more preferably 1-2.5 μg / mg, such as 1.2 μg / mg, 1.4 μg / mg, 1.5 μg / mg, 1.6 μg / mg, 1.8 μg / mg, 2 μg / mg, 2.2 μg / mg, or 2.4 μg / mg.

[0036] In another preferred embodiment, step S1 includes the step of adjusting the mechanical properties of the dual-network hydrogel by adjusting the concentration and / or degree of substitution of the (meth)acrylamide gelatin in the aqueous precursor solution.

[0037] In another preferred embodiment, the mechanical property is selected from: storage modulus, loss modulus, stress relaxation time, Young's modulus, or a combination thereof.

[0038] In another preferred embodiment, the energy storage modulus can be adjusted within the range of 20 Pa to 1000 Pa, preferably within the range of 50 Pa to 500 Pa.

[0039] In another preferred embodiment, the loss modulus can be adjusted in the range of 20 Pa to 10 kPa, preferably in the range of 50 Pa to 5 kPa or 50 Pa to 3 kPa.

[0040] In another preferred embodiment, the stress relaxation time can be adjusted in the range of 0.1 min to 5 min, preferably in the range of 0.3 min to 3 min, and more preferably in the range of 0.5 min to 2 min.

[0041] In another preferred embodiment, the Young's modulus can be adjusted in the range of 0.1 kPa to 50 kPa, preferably in the range of 0.2 kPa to 15 kPa, more preferably in the range of 0.5 kPa to 10 kPa, and even more preferably in the range of 0.5 kPa to 2.5 kPa.

[0042] In another preferred embodiment, reducing the concentration of (meth)acrylamide gelatin in the aqueous precursor solution reduces the storage modulus, and vice versa.

[0043] In another preferred embodiment, reducing the concentration of (meth)acrylamide gelatin in the aqueous precursor solution and / or increasing the average degree of substitution of (meth)acrylamide gelatin reduces the loss modulus, and vice versa.

[0044] In another preferred embodiment, fixing the concentration of (meth)acrylamide gelatin in the aqueous precursor solution reduces the average degree of substitution of (meth)acrylamide gelatin, thereby reducing the stress relaxation time, and vice versa.

[0045] In another preferred embodiment, fixing the average degree of substitution of (meth)acrylamide gelatin reduces the concentration of (meth)acrylamide gelatin in the aqueous precursor solution, thereby reducing Young's modulus, and vice versa.

[0046] In another preferred embodiment, the decellularized matrix is ​​derived from neural tissues such as the spinal cord, brain, and sciatic nerve, or from non-neural tissues such as the submucosa of the small intestine, skin, and skeletal muscle, preferably neural tissues such as the spinal cord.

[0047] In another preferred embodiment, the decellularized matrix is ​​prepared by the following method: Step i: Cell membrane disruption: Fresh spinal cord tissue is cut into pieces and placed in a low-temperature freezer (e.g., -60℃ to -80℃) for repeated freeze-thaw cycles to disrupt the cell membrane; Step ii: Chemically enzymatically hydrolyze the product from step i to obtain the chemically enzymatically hydrolyzed product; Step iii: The chemical enzymatic hydrolysis product obtained in step ii is added at a mass fraction of 0.1-1% w / w (preferably 0.2-0.8%, more preferably 0.4-0.6%) to a 0.008-0.015 mol / L HCl solution (preferably 0.01-0.012 mol / L) containing 0.3-1 mg / mL (preferably 0.4-0.8%, more preferably 0.4-0.6%) pepsin. The mixture is continuously stirred and hydrolyzed at 20-38°C (preferably 25-37°C) until completely dissolved. The pepsin hydrolysis product is then freeze-dried to obtain the decellularized matrix.

[0048] In another preferred embodiment, the chemical enzymatic hydrolysis includes the step of repeating the following treatment steps on the tissue obtained in step i for 2-4 cycles: Water treatment for 4-8 hours, followed by treatment with 2-4 v / v% Triton X-100 solution for 5-8 hours; Water treatment for 15-30 min × 3 times, followed by treatment with 0.03-0.05 g / mL sodium deoxycholate solution for 8-12 h; Treat with water for 15-30 minutes × 3 times, then treat with DNase for 3-5 hours; The chemical enzymatic hydrolysis product is then obtained by freeze-drying after the following cleaning process: Water treatment for 15-30 min × 3 times, 3-5 v / v% ethanol treatment for 3-5 h, and water treatment for 15-30 min × 3 times; All processing steps were carried out in a constant temperature shaker at 37±2℃ with a rotation speed set at 100-200 rpm. After each reagent was processed, the liquid was discarded before proceeding to the next operation.

[0049] In another preferred embodiment, the water is selected from the group consisting of: ultrapure water, double-distilled water, and deionized water.

[0050] In another preferred embodiment, the chemical enzymatic decomposition of cells includes: treating the tissue obtained in step i after cell membrane disruption according to the following order, treatment reagents, and treatment time: All steps were carried out in a constant temperature shaker at 37±0.5℃, with the speed set at 100-200 rpm. After each reagent was processed, the liquid was discarded before proceeding to the next operation. First cycle: Ultrapure water (4-8h), 2-4v / v% Triton X-100 (5-8h), ultrapure water (15-30min×3), 0.03-0.05g / mL SDC (8-12h), ultrapure water (15-30min×3), 3500-4500 kunits / ml DNase (3-5h); Second cycle: Ultrapure water (15-30 min × 3), 2-4 v / v% Triton X-100 (5-8 h), ultrapure water (15-30 min × 3), 0.03-0.05 g / mL SDC (8-12 h), ultrapure water (15-30 min × 3), 3500-4500 kunits / ml DNase (3-5 h); Third cycle: Ultrapure water (15-20 min × 3), 2-4 v / v% Triton X-100 (5-8 h), ultrapure water (15-30 min × 3), 0.03-0.05 g / mL SDC (8-12 h), ultrapure water (15-30 min × 3), 3500-4500 kunits / ml DNA enzyme (3-5 h), ultrapure water (15-30 min × 3), 3-5 v / v% ethanol (3-5 h), ultrapure water (15-30 min × 3).

[0051] In another preferred embodiment, the first cycle consisted of: ultrapure water (6 h), 3 v / v% Triton X-100 (6 h), ultrapure water (20 min × 3), 0.04 g / mL SDC (10 h), ultrapure water (20 min × 3), and 4000 kunits / ml DNA enzyme (4 h).

[0052] In another preferred embodiment, the second cycle consists of: ultrapure water (20 min × 3), 3 v / v% Triton X-100 (6 h), ultrapure water (20 min × 3), 0.04 g / mL SDC (10 h), ultrapure water (20 min × 3), and 4000 DNA enzymes (4 h).

[0053] In another preferred embodiment, the third cycle consists of: ultrapure water (20 min × 3), 3 v / v% Triton X-100 (6 h), ultrapure water (20 min × 3), 0.04 g / mL SDC (10 h), ultrapure water (20 min × 3), 4000 DNA enzyme (4 h), ultrapure water (20 min × 3), 4 v / v% ethanol (4 h), and ultrapure water (20 min × 3).

[0054] In another preferred embodiment, the residual DNA content in the decellularized matrix is ​​less than 50 ng / mg, on a dry weight basis.

[0055] In a second aspect, the present invention provides a decellularized matrix dual-network hydrogel prepared by the preparation method described in the first aspect of the present invention.

[0056] In another preferred embodiment, the dual-network hydrogel is in the form of microspheres.

[0057] In this invention, the dual-network hydrogel has an interpenetrating network structure, wherein the first network is a physical network formed by the self-assembly of decellularized matrix (without the need for additional crosslinking agents), and the second network is a chemical network formed by photocrosslinking of methacryloyl polymer.

[0058] In another preferred embodiment, the dual-network hydrogel is injectable.

[0059] In another preferred embodiment, the average pore size (second network) of the dual-network hydrogel is 10-30 μm, preferably 15-25 μm.

[0060] In another preferred embodiment, the dual-network hydrogel has a low stress relaxation time (e.g., ≤1 min), preferably 0.1-1 min, more preferably 0.5-0.8 min.

[0061] In another preferred embodiment, the Young's modulus of the dual-network hydrogel is 0.5 kPa-2.5 kPa, preferably 0.6 kPa-2 kPa.

[0062] In another preferred embodiment, the loss modulus of the dual-network hydrogel is 10 Pa to 100 Pa, preferably 10 to 50 Pa.

[0063] In another preferred embodiment, the storage modulus of the dual-network hydrogel is 20 Pa to 1000 Pa, preferably 100 Pa to 300 Pa.

[0064] A third aspect of the present invention provides a cell culture substrate or tissue engineering scaffold comprising: a decellularized matrix dual-network hydrogel according to a second aspect of the present invention; the decellularized matrix dual-network hydrogel having a covalently modified heparin layer and loaded with cytokines having heparin binding domains.

[0065] In another preferred embodiment, the cytokine is selected from one or more of the following: brain-derived neurotrophic factor (BDNF), nerve growth factor (NGF), vascular endothelial growth factor (VEGF), fibroblast growth factor (FGF), hepatocyte growth factor (HGF), epidermal growth factor (EGF), stromal cell-derived factor-1 (SDF-1), and neurotrophic factor-3 (NT-3).

[0066] In another preferred embodiment, the dual-network hydrogel is in the form of microspheres. In yet another preferred embodiment, the average diameter of the microspheres is 50-400 μm, preferably 150-300 μm.

[0067] In another preferred embodiment, the tissue-engineered scaffold is injectable.

[0068] In a fourth aspect, the present invention provides the use of the dual-network hydrogel described in the second aspect of the present invention in the preparation of tissue engineering scaffolds, drug carriers and / or cell culture substrates.

[0069] In another preferred embodiment, the tissue-engineered scaffold is a neural-engineered scaffold.

[0070] In another preferred embodiment, the tissue is selected from the group consisting of: spinal cord, brain, peripheral nerves, and periosteum.

[0071] In another preferred embodiment, the cells are selected from the group consisting of neural stem cells, astrocytes, microglia, neurons, mesenchymal stem cells, and Schwann cells.

[0072] It should be understood that, within the scope of this invention, the above-described technical features of this invention and the technical features specifically described below (such as in the embodiments) can be combined with each other to form new or preferred technical solutions. Due to space limitations, they will not be described in detail here. Attached Figure Description

[0073] Figure 1This is a comparison chart of the residual DNA content of porcine spinal cord tissue before and after decellularization treatment (dECM) in Example 1 of the present invention. The data in the chart shows that after treatment by the method of the present invention, the DNA content is significantly reduced from about 480 ng / mg in the native tissue to below 50 ng / mg (about 45 ng / mg), which proves the high efficiency and safety of the decellularization process.

[0074] Figure 2 This is the preparation process and basic characterization diagram of the dual-network hydrogel microspheres in Embodiment 2 of the present invention; the upper part is a schematic diagram of the principle of microfluidic chip preparation of microspheres, the upper right part is an optical microscope photograph of the microspheres, and the lower part is a photograph of the microfluidic chip and a histogram of the particle size distribution of the microspheres; the histogram shows that the microspheres have a uniform particle size distribution, which is normally distributed, with an average diameter of about 210 μm.

[0075] Figure 3 These are scanning electron microscope (SEM) morphology images of the dual-network hydrogel microspheres prepared in Example 2 of this invention; the left image shows the overall morphology of the microspheres, while the middle and right images are magnified views of the local areas, showing that the surface and interior of the microspheres have a dense porous network structure that penetrates each other, confirming the successful construction of the dual-network structure.

[0076] Figure 4 This is a scanning electron microscope (SEM) morphology characterization image of the dual-network hydrogel prepared in Example 3 of this invention.

[0077] Figure 5 Figure 1 shows the mechanical property test results of the dual-network hydrogels with different GelMA concentration formulations (groups A, B, C, and D) in Example 3 of this invention; where Figure A is the rheological curve of storage modulus (G') as a function of frequency, Figure B is the rheological curve of loss modulus (G'') as a function of frequency, Figure C is a comparison of Young's modulus of each group of samples, and Figure D is a comparison of stress relaxation time of each group of samples.

[0078] Figure 6 This is a comparison chart of the quantitative detection of heparin content (glycosaminoglycans GAGs) on the surface of ungrafted heparin microspheres (MS) and heparin-grafted microspheres (Hep-MS) in Example 4 of the present invention.

[0079] Figure 7 This is a statistical chart showing the cell viability (CCK-8 assay) of cells cultured in different microsphere (MS, Hep-MS) extracts for 1 day (left figure) and 3 days (right figure) in Example 5 of the present invention.

[0080] Figure 8This is a microscopic image (cytoskeleton staining) of mouse hippocampal neurons (HT-22) cultured on the surface of dual-network hydrogel microspheres in Example 5 of this invention; in the image, blue represents DAPI-labeled cell nuclei, and green represents F-actin labeled with phalloidin.

[0081] Figure 9 This is a scanning electron microscope (SEM) morphology characterization of heparin-grafted microspheres (Hep-MS) after 7 days of in vitro degradation, showing that the second network degrades first.

[0082] Figure 10 This demonstrates the retention quality of the dual-network hydrogel of the present invention after degradation over time.

[0083] Figure 11 The image shows the morphology of the material obtained by first performing photocrosslinking of the second network and then thermal crosslinking of the first network, as indicated by a scanning electron microscope (SEM). It shows that the first network cannot be formed.

[0084] Figure 12 This is a scanning electron microscope (SEM) morphology characterization of the fibrous network of the first network of decellularized matrix after gelation.

[0085] Figure 13 This shows the gel formation of the decellularized matrix solutions in Comparative Example 2, which were undigested (right side) and digested (left side).

[0086] Figure 14 This demonstrates the effect of whether or not a decellularized matrix first network is formed on the mechanical properties of GelMA gel. Detailed Implementation

[0087] Through extensive and in-depth research, including numerous screenings and tests, the inventors have developed a decellularized matrix dual-network hydrogel, its preparation method, and its applications. The hydrogel carrier provides a nerve-specific microenvironment through a first network of decellularized matrix, achieves independent regulation of mechanical properties through a chemically cross-linked second network, and can further achieve long-term sustained release of neurotrophic factors through covalently grafted heparin. This invention, through specific component combinations and preparation methods, solves the problem of existing microcarriers that struggle to simultaneously achieve bioactivity, mechanical compatibility, degradation performance, and drug sustained-release function. Based on this, the inventors completed this invention.

[0088] the term Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention pertains.

[0089] As used herein, the terms “containing” or “including (comprise)” can be open-ended, semi-closed, or closed. In other words, the terms also include “consistently made of” or “made of”.

[0090] As used in this article, the terms "room temperature" or "normal temperature" refer to a temperature between 4 and 40 degrees Celsius. o C, preferably, 25±5 o C.

[0091] Dual-network hydrogels and their preparation This invention provides a dual-network hydrogel based on a decellularized matrix, comprising a first network formed by the thermal gelation of the decellularized matrix and a second network formed by (meth)acrylamide gelatin. The two networks interpenetrate to form the dual-network hydrogel. In other words, the first and second networks in the dual-network hydrogel of this invention are inseparable, and the hydrogel is substantially homogeneous overall. Compared to polymer-encapsulated decellularized matrix, this interpenetrating network eliminates the shielding effect of the polymer shell on internal biological activity, allowing cells to simultaneously acquire RGD adhesion sites (provided by amino-containing methacrylamide polymers) and dECM neural induction signals upon contact with the material surface. That is, the dECM component can exert its effects immediately without waiting for the external material to degrade. Furthermore, the mechanism by which the second network (GelMA) degrades before the first network (dECM) creates microporous channels in situ within the hydrogel, actively reducing physical steric hindrance and guiding cells to infiltrate deeper into the scaffold.

[0092] Specifically, the dual-network hydrogel of the present invention can be prepared by a method comprising the following steps: S1. Preparation of aqueous precursor solution: Dissolve decellularized matrix (dECM) in water under acidic conditions, then mix with an aqueous solution of (meth)acrylamide gelatin and photoinitiator, and adjust the pH to neutral at low temperature to obtain an aqueous precursor solution; S2. Primary thermal gelation: The aqueous precursor solution is heated to induce primary thermal gelation of the decellularized matrix, forming a mixture containing a first network gel; S3. Photocrosslinking and curing: Applying light of a predetermined wavelength to the mixture containing the first network gel initiates a polymerization reaction of (meth)acrylamide gelatin to form a second network, thereby obtaining an interpenetrating double network hydrogel.

[0093] In particular, in this invention, the second network cannot be cured before the first network because the thermal gelation of the decellularized matrix cannot be achieved after the second network is formed. Preferably, the photocuring of the second network is performed after the first network has been fully formed.

[0094] In another preferred embodiment, the method further includes: S4. Functional modification: covalently grafting heparin molecules onto the surface of the dual-network hydrogel via an amidation reaction.

[0095] In this invention, the (meth)acrylamide gelatin is selected from methacrylamide gelatin and acrylamide gelatin, preferably methacrylamide gelatin.

[0096] In another preferred embodiment, the concentration of decellularized matrix in the aqueous precursor solution is 5-50 mg / mL (0.5-5.0 w / v%), preferably 7.5-20 mg / mL, and more preferably 10-15 mg / mL. Within this concentration range, dECM can provide sufficient bioactive sites and extracellular matrix microenvironment, while avoiding a surge in solution viscosity due to excessively high concentrations, ensuring good monodispersity of the microfluidically generated droplets.

[0097] In another preferred embodiment, the concentration of (meth)acrylamide gelatin in the aqueous precursor solution is 1-20 w / v%, preferably 2-10 w / v%, more preferably 3-6 w / v%. The second network formed by the (meth)acrylamide gelatin mainly provides mechanical support for the carrier. By adjusting the concentration and / or degree of substitution of this component, the Young's modulus of the carrier can be precisely controlled within the range of 0.1 kPa to 50 kPa (especially within the range of 0.5 kPa to 15 kPa) to match the physiological and mechanical properties of different target tissues. The aqueous precursor solution needs to be stored at a low temperature, preferably 0-10°C, more preferably 0-4°C.

[0098] In another preferred embodiment, the photoinitiator may be any group of photoinitiators suitable for methacryl or acrylyl groups in the art, including (but not limited to): lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP), lithium 2,4,6-trimethylbenzoyl-diphenylphosphine oxide (Li-TPO), camphorquinone (CQ), riboflavin, or one or more thereof. The wavelength, light intensity, and time of the initiation reaction may be selected according to the photoinitiator.

[0099] In another preferred embodiment, the aqueous solution of (meth)acrylamide gelatin and photoinitiator can be (meth)acrylamide gelatin dissolved in phosphate-buffered saline (PBS) containing 0.1-0.4% (w / v) (preferably 0.2-0.3%) of photoinitiator.

[0100] In another preferred embodiment, the (meth)acrylamide gelatin has an average molecular weight of 50 kDa to 100 kDa.

[0101] In another preferred embodiment, the (meth)acrylamide gelatin (GelMA) is prepared by reacting type A gelatin with methacrylic anhydride or acrylic anhydride. The type A gelatin may be derived from pigskin, bovine bones, or bovine hide.

[0102] In another preferred embodiment, the Bloom value of the methacrylamide gelatin is 300-350, preferably 310-330, and more preferably 320-325.

[0103] In another preferred embodiment, the degree of substitution of the (meth)acrylamide gelatin is selected from one or more of 5-95%, and can be adjusted by compounding as needed. As used in this invention, the degree of substitution is the molar percentage of the number of amino groups on the (meth)acrylamide gelatin molecular chain that have undergone methacrylation or acrylation modification relative to the initial total number of amino groups on its molecular chain.

[0104] Preferably, the (meth)acrylamide gelatin is a mixture of (meth)acrylamide gelatin with low and high degrees of substitution. This is advantageous for preparing a dual-network hydrogel with a low stress relaxation time. For example, the (meth)acrylamide gelatin is a mixture of (meth)acrylamide gelatin with substitution degrees of 10%-20% and 85%-95%.

[0105] In another preferred embodiment, step S2 includes the following steps: using the aqueous precursor solution as the dispersed phase and an oily liquid containing a surfactant as the continuous phase, droplet generation is performed in a microfluidic chip, and primary thermal gelation of the decellularized matrix is ​​induced at the rear end of the chip by heating to form a mixture microsphere containing a first network gel.

[0106] In another preferred embodiment, the oily liquid is selected from one or more of liquid paraffin, mineral oil, vegetable oil, and fluorinated oil.

[0107] In another preferred embodiment, the surfactant is selected from one or more of Span, Tween, and fluorinated surfactants, preferably Span 80.

[0108] In another preferred embodiment, the amount of surfactant added in the continuous phase is 0.1-10.0 w / v%, preferably 2-8 w / v%, more preferably 4-6 w / v.

[0109] In another preferred embodiment, in step S2, the heating temperature is 30℃-50℃, 35-45℃, and preferably 37-40℃. When the step is performed in a microfluidic chip, the temperature can be increased by controlling the temperature at the end of the microfluidic chip (e.g., by placing it on a temperature-controlled operating table). This invention utilizes physiological temperature ranges to induce the self-assembly of collagen fibers, and uses thermo-induced physical gelation to improve the morphological stability of droplets before photocuring, which can also prevent droplet fusion.

[0110] The initial thermogelation process of this invention should typically be carried out under low disturbance or static conditions, as stirring or other operations may prevent the components in the decellularized matrix from self-assembling. The initial thermogelation time is usually 5-30 minutes, preferably 10-20 minutes.

[0111] In another preferred embodiment, during step S2, the flow rate ratio of the dispersed phase to the continuous phase during droplet generation is 1:(1-20), preferably 1:(2-10), and more preferably 1:(5-8). By using a microfluidic chip to prepare a W / O emulsion, and adjusting the flow rate ratio, microspheres with a uniform particle size distribution and an adjustable diameter between 50-400 μm can be obtained.

[0112] In another preferred embodiment, the diameter of the microspheres is 100-300 μm, preferably 150-250 μm.

[0113] In another preferred embodiment, the average pore size of the dual-network hydrogel is 10-30 μm, preferably 15-25 μm.

[0114] In another preferred embodiment, in step S4, the amidation reaction involves reacting the carboxyl groups in the heparin molecule with the amino groups (including those on the decellularized matrix and (meth)acrylamide gelatin) on the surface of the dual-network hydrogel. Preferably, the reaction is carried out in the presence of a coupling agent (such as EDC / NHS). Compared to physical adsorption, covalent binding can significantly reduce the "burst release effect" of the loading factor, achieving long-lasting sustained release.

[0115] In this invention, there are no special requirements regarding the tissue source of the decellularized matrix; the decellularized matrix of the tissue to be repaired can be selected as needed. For example, the decellularized matrix may be derived from neural tissues such as the spinal cord, brain, and sciatic nerve, or from non-neural tissues such as the submucosa of the small intestine, skin, and skeletal muscle.

[0116] In particular, the mechanical properties and degradation properties of the dual-network hydrogel of the present invention are particularly suitable for decellularized matrices derived from nerve tissue, and can well match the dynamic needs of the nerve tissue repair cycle during application.

[0117] Typically, the decellularized matrix can be prepared by methods known in the art. Preferably, the residual DNA content in the decellularized matrix is ​​less than 50 ng / mg, on a dry weight basis.

[0118] In another preferred embodiment, the decellularized matrix is ​​prepared by the following method: Step i: Cell membrane disruption: Take fresh tissue, cut it into pieces, and place it in a low-temperature freezer (e.g., -60℃ to -80℃) for repeated freeze-thaw cycles to disrupt the cell membrane; Step ii: The product of step i is subjected to chemical enzymatic hydrolysis to obtain the chemical enzymatic hydrolysis product; The chemical enzymatic hydrolysis product obtained in step ii is added to an acidic aqueous solution containing pepsin, and the mixture is stirred and enzymatically hydrolyzed until completely dissolved. The enzymatic hydrolysis product is then freeze-dried to obtain the decellularized matrix.

[0119] In another preferred embodiment, step iii: The chemical enzymatic hydrolysis product obtained in step ii is added at a mass fraction of 0.1-1% w / w (preferably 0.2-0.8%, more preferably 0.4-0.6%) to a 0.008-0.015 mol / L HCl solution (preferably 0.01-0.012 mol / L) containing 0.3-1 mg / mL (preferably 0.4-0.8%, more preferably 0.4-0.6%) pepsin. The mixture is continuously stirred and hydrolyzed at 20-38°C (preferably 25-37°C) until completely dissolved. The pepsin hydrolysis product is then freeze-dried to obtain the decellularized matrix.

[0120] Preferably, the decellularized matrix can be derived from humans or non-human mammals, such as rats, mice, monkeys, and pigs. More preferably, it is derived from an allogeneic organism.

[0121] Methods for controlling mechanical properties Specifically, in step S1, the mechanical properties of the dual-network hydrogel can be adjusted by regulating the concentration and / or degree of substitution of the (meth)acrylamide gelatin in the aqueous precursor solution.

[0122] In another preferred embodiment, the mechanical property is selected from: storage modulus, loss modulus, stress relaxation time, Young's modulus, or a combination thereof.

[0123] In another preferred embodiment, the energy storage modulus can be adjusted within the range of 20 Pa to 1000 Pa, preferably within the range of 50 Pa to 500 Pa.

[0124] In another preferred embodiment, the loss modulus can be adjusted in the range of 20 Pa to 10 kPa, preferably in the range of 50 Pa to 5 kPa or 50 Pa to 3 kPa.

[0125] In another preferred embodiment, the stress relaxation time can be adjusted in the range of 0.1 min to 5 min, preferably in the range of 0.3 min to 3 min, and more preferably in the range of 0.5 min to 2 min.

[0126] In another preferred embodiment, the Young's modulus can be adjusted in the range of 0.1 kPa to 50 kPa, preferably in the range of 0.2 kPa to 15 kPa, more preferably in the range of 0.5 kPa to 10 kPa, and most preferably in the range of 0.5 kPa to 2.5 kPa.

[0127] In another preferred embodiment, reducing the concentration of (meth)acrylamide gelatin in the aqueous precursor solution reduces the storage modulus, and vice versa.

[0128] In another preferred embodiment, reducing the concentration of (meth)acrylamide gelatin in the aqueous precursor solution and / or increasing the average degree of substitution of (meth)acrylamide gelatin reduces the loss modulus, and vice versa.

[0129] In another preferred embodiment, fixing the concentration of (meth)acrylamide gelatin in the aqueous precursor solution reduces the average degree of substitution of (meth)acrylamide gelatin, thereby reducing the stress relaxation time, and vice versa.

[0130] In another preferred embodiment, fixing the average degree of substitution of (meth)acrylamide gelatin reduces the concentration of (meth)acrylamide gelatin in the aqueous precursor solution, thereby reducing Young's modulus, and vice versa.

[0131] Preferably, the average degree of substitution of (meth)acrylamide gelatin can be achieved by using (meth)acrylamide gelatin with a lower (but substantially consistent) degree of substitution. More preferably, a mixture of (meth)acrylamide gelatin with low and high degrees of substitution can be used, where the degrees of substitution can differ significantly (e.g., by 50% or more). By adjusting the mixing ratio of the two, the desired average degree of substitution of (meth)acrylamide gelatin can be precisely obtained. As described, the (meth)acrylamide gelatin is a mixture of (meth)acrylamide gelatin with a substitution degree of 10%-20% and (meth)acrylamide gelatin with a substitution degree of 85%-95%. This is advantageous for preparing dual-network hydrogels with low stress relaxation times.

[0132] The method of this invention allows for flexible control of the properties of the resulting dual-network hydrogel, and surprisingly, the method can independently control the viscoelastic properties and mechanical strength of the dual-network hydrogel. For example, it can yield materials that simultaneously possess low stress relaxation time and low Young's modulus, properties that are typically difficult to prepare in the art.

[0133] application The dual-network hydrogel described in this invention possesses the bioactivity of the decellularized matrix itself, tunable mechanical strength, good biocompatibility, biodegradability, and drug loading / sustained release capabilities. It can serve as a platform material with significant application prospects in tissue repair and regenerative medicine, particularly in neural regeneration and stem cell proliferation and induced differentiation.

[0134] Therefore, the aforementioned dual-network hydrogel can be used as a tissue engineering scaffold, drug carrier, and / or cell culture carrier, etc.

[0135] The present invention provides a cell culture substrate or tissue engineering scaffold comprising: a decellularized matrix dual-network hydrogel prepared by the method described in the first aspect of the present invention; the dual-network hydrogel having a covalently modified heparin layer and loaded with cytokines having heparin binding domains.

[0136] In another preferred embodiment, the cytokine is selected from one or more of the following: brain-derived neurotrophic factor (BDNF), nerve growth factor (NGF), vascular endothelial growth factor (VEGF), fibroblast growth factor (FGF), hepatocyte growth factor (HGF), epidermal growth factor (EGF), stromal cell-derived factor-1 (SDF-1), and neurotrophic factor-3 (NT-3).

[0137] In another preferred embodiment, the dual-network hydrogel is in the form of microspheres. In yet another preferred embodiment, the average diameter of the microspheres is 50-400 μm, preferably 150-300 μm.

[0138] In another preferred embodiment, the tissue-engineered scaffold is injectable.

[0139] In another preferred embodiment, the tissue-engineered scaffold is a neural-engineered scaffold.

[0140] The main advantages of this invention include: 1. By constructing a dual-network interpenetrating structure, the decoupled regulation of the biological function and physical properties of microspheres is achieved. In this invention, the first network, composed of decellularized matrix (dECM), is mainly responsible for providing tissue-specific bio-inducing signals, thus preserving its biological activity. The second network, composed of chemically cross-linked polymers, serves as structural support, allowing for independent and precise setting of the carrier's mechanical strength, viscoelasticity, and degradation rate by adjusting parameters such as precursor concentration or cross-linking density. This method overcomes the technical challenge of the mutual constraint between biological activity and physical properties in single-component materials.

[0141] 2. This invention provides a multimodal, guiding microenvironment for regulating cell behavior. It integrates two different types of regulatory signals. First, the dECM network provides complex biochemical and topological signals that effectively support cell growth and differentiation. Second, the tunable mechanical properties of the second network provide physical signal input for influencing cell behavior (including maintaining stemness or guiding differentiation) using the principle of mechanical transduction. This organic combination of two signaling modalities makes more precise and strategic targeted regulation of stem cell fate possible.

[0142] 3. By covalently binding heparin, microspheres are transformed from passive scaffolds into active biological signal regulation platforms. The heparin portion utilizes its specific affinity for bioactive molecules with heparin-binding domains (such as various growth factors) to achieve effective capture, local enrichment, and bioactivity protection of these molecules. Therefore, this invention can construct a controllable reservoir of signaling molecules and achieve precise spatiotemporal delivery within the microenvironment, which is difficult to achieve using simple physical encapsulation methods in the prior art.

[0143] The present invention will be further illustrated below with reference to specific embodiments. It should be understood that these embodiments are for illustrative purposes only and are not intended to limit the scope of the invention. Experimental methods in the following embodiments, unless otherwise specified, are generally performed under conventional conditions or as recommended by the manufacturer. Percentages and parts are by weight unless otherwise stated.

[0144] Explanation of Terminology Abbreviations dECM: Decellularized matrix of the spinal cord GelMA: Methacrylated gelatin LAP: Lithium phenyl-2,4,6-trimethylbenzoylphosphonate (photoinitiator) EDC: 1-(3-Dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride NHS: N-hydroxysuccinimide HT-22: Mouse hippocampal neurons G': Energy storage modulus G'': Loss modulus E: Young's modulus.

[0145] Example 1: Preparation and characterization of decellularized spinal cord matrix (dECM) Preparation of decellularized porcine spinal cord matrix dECM (1) Take fresh pig spinal cord tissue, remove the pia mater and blood vessels, cut it into pieces and freeze it in a -80℃ freezer. Repeat the freeze-thaw cycle three times to break the cell membrane.

[0146] (2) Chemical enzymatic decomposition of cells: List the treatment reagents and treatment times in order: First cycle: Place the cleaned spinal cord tissue block in a wide-mouth bottle. All elution steps are performed in a constant-temperature shaker at 37°C with a rotation speed set at 140 rpm.

[0147] First, wash with ultrapure water for 6 hours and discard the waste liquid; then add 3% (v / v) Triton X-100 solution and wash with water for 6 hours; then discard the waste liquid and wash with ultrapure water 3 times, 20 min each time; next, add 0.04 g / mL sodium deoxycholate (SDC) solution and wash with water for 10 hours; wash with ultrapure water 3 times again, 20 min each time; finally, add 4000 kunits / ml DNase enzyme solution and digest with water for 4 hours. Second cycle: Ultrapure water (20 min × 3), 3 v / v% Triton X-100 (6 h), ultrapure water (20 min × 3), 0.04 g / mL SDC (10 h), ultrapure water (20 min × 3), 4000 kunits / ml DNase (4 h); Third cycle: ultrapure water (20 min × 3), 3 v / v% Triton X-100 (6 h), ultrapure water (20 min × 3), 0.04 g / mL SDC (10 h), ultrapure water (20 min × 3), 4000 kunits / ml DNase (4 h), ultrapure water (20 min × 3), 4 v / v% ethanol (4 h), ultrapure water (20 min × 3); then freeze-dry to obtain chemically hydrolyzed dECM.

[0148] DNA residual detection Residual DNA was extracted from the chemically enzymatically hydrolyzed dECM using a DNA extraction kit, and the DNA content was detected using picogreen. The results showed that the residual DNA content in the decellularized dECM was less than 50 ng / mg (dry weight). Figure 1 It meets the safety standards for biomedical materials, proving that the decellularization process is effective.

[0149] (3) Enzymatic hydrolysis by proteases Weigh out the lyophilized powder of decellularized spinal cord matrix obtained by chemical enzymatic hydrolysis, and add it to a 0.01 mol / L HCl solution containing 0.5 mg / mL pepsin at a mass fraction of 0.5%. Continue stirring at room temperature for 48 hours until completely dissolved. Freeze-dry the digestion product to obtain purified dECM enzymatically hydrolyzed lyophilized powder.

[0150] Example 2: Preparation of dECM / GelMA dual-network hydrogel microspheres In this embodiment, a dual-network microsphere substrate is prepared using microfluidic technology. The specific steps are as follows: (1) Microfluidic chip preparation Microfluidic chips with flow focusing structures were fabricated using polydimethylsiloxane (PDMS). The width and depth of the dispersed phase channel, continuous phase channel, and main channel were all 200 μm. Figure 2 ).

[0151] (2) Preparation of aqueous and oil phases Aqueous phase: The dECM enzymatically hydrolyzed lyophilized powder prepared in Example 1 was reconstituted in 0.01 M HCl solution to prepare a 1% (w / w) dECM acidic stock solution. Separately, 4% (w / v) methacrylamide gelatin (GelMA) (95% degree of substitution, 320 g bloom porcine skin-derived type A gelatin) was dissolved in phosphate-buffered saline (PBS) containing 0.25% (w / v) photoinitiator (LAP) and stirred in a 40°C water bath to prepare a GelMA prepolymer solution. The above dECM acidic stock solution and GelMA prepolymer solution were mixed in a predetermined ratio, and the pH was carefully adjusted to 7.4 (neutral) using NaOH solution. The mixture was then stored at low temperature (4°C) as the aqueous phase for later use.

[0152] Oil phase: Dissolve 5% (w / v) of Span 80 in liquid paraffin oil and mix thoroughly.

[0153] (3) Microsphere formation and solidification The aqueous phase (4 μL / min) and oil phase (24 μL / min) were pumped separately into the microfluidic chip. Figure 2 The flow rate ratio was controlled at 1:6. The chip and the primary collection device were placed on a 37°C constant-temperature heating stage. The droplet remained in the primary collection device for about 10 minutes, and the temperature induced the physical thermogelation of the dECM components within the droplet (formation of the first network). Subsequently, the collected liquid was passed through a blue light region with a wavelength of 405 nm (light intensity of 30 mW / cm²). 2 Irradiation for 40 seconds initiated GelMA polymerization (second network formation). The product was demulsified with a demulsifier (1% Twen80), washed three times with PBS, and dECM / GelMA dual-network microspheres were obtained.

[0154] (4) Morphological characteristics Optical microscopy observation showed ( Figure 2 ) and Scanning Electron Microscopy (SEM) Figure 3 The microspheres exhibit a regular spherical shape, good monodispersity, an average particle size of approximately 210 μm, and an average pore size of 18.64 ± 4.45 μm. The microspheres also display a dense, interpenetrating porous network structure, confirming the successful construction of the dECM and GelMA dual network.

[0155] Example 3: Screening and Characterization of Microsphere Formulations with Controllable Mechanical Strength To verify the mechanical tunability of the dual-network hydrogel, this example prepared lyophilized powder containing the same concentration of dECM enzyme hydrolysis (prepared in Example 1, dissolved in 0.01 M HCl solution) but with different GelMA formulations (dissolved in phosphate-buffered saline (PBS) containing 0.25% (w / v) photoinitiator (LAP)). The preparation method involved thoroughly mixing the aqueous precursor solution, injecting it into a mold, heating to 37°C and maintaining the temperature for 10 min. After the first network formed, 405 nm blue light (30 mW / cm²) was used. 2 The gel was irradiated for 40 seconds to cure it using light, resulting in large cylindrical hydrogels (20 mm in diameter and 1.5 mm in height). Typical micrographs are shown below. Figure 4 Mechanical characterization was performed using a rheometer to simulate the matrix hardness of the corresponding microspheres.

[0156] (1) Experimental Grouping Group A: 4% (w / v) GelMA (95% degree of substitution) + 1% (w / v) dECM enzymatically hydrolyzed lyophilized powder; Group B: 1.6% (w / v) GelMA (95% degree of substitution), 2.3% (w / v) GelMA (15% degree of substitution), + 1% (w / v) dECM enzymatically hydrolyzed lyophilized powder; Group C: 8% (w / v) GelMA (95% degree of substitution) + 1% (w / v) dECM enzymatically hydrolyzed lyophilized powder; Group D: 5% (w / v) GelMA (95% degree of substitution), 4% (w / v) GelMA (15% degree of substitution), + 1% (w / v) dECM enzymatically hydrolyzed lyophilized powder; (2) Rheological testing Frequency scanning was performed at 37℃, and the storage modulus (G') and loss modulus (G'') were recorded. The test results show that the rheological properties can be precisely controlled (50 Pa-500 Pa) by adjusting the concentration and degree of substitution of GelMA.

[0157] An instantaneous shear strain is applied within the linear viscoelastic region of the sample, and the stress decay over time is recorded to obtain its stress relaxation time.

[0158] The results confirmed that the rheological behavior of hydrogels can be independently decoupled and regulated by mixing GelMA with different degrees of substitution, taking advantage of the differences in free segment length and crosslinking density within the system. Furthermore, the introduction of the first network can result in a faster stress relaxation time compared to traditional GelMA gels, which is beneficial for cell growth and ingrowth.

[0159] (3) Young's modulus test Compression tests were conducted on the specimens at a constant strain rate of 0.02 mm / s at 37°C. Young's modulus was calculated based on the first 10% deformation of the hydrogel. The results confirmed that the mechanical strength of the hydrogel (0.5 kPa–2.5 kPa) can be independently controlled by adjusting the concentration and degree of substitution of GelMA. Figure 5 The data show that the Young's modulus of the carrier in the embodiments can be adjusted in the range of about 600 Pa to 2.5 kPa, and the differences between groups are significant (**** p<0.0001).

[0160] In summary, while groups A and B exhibit similar storage and loss moduli, their stress relaxation times differ significantly. Conversely, groups A and C show similar stress relaxation behaviors, but their Young's moduli are drastically different. This result strongly confirms that by mixing GelMA with different degrees of substitution and utilizing the differences in free segment length and crosslinking density within the system, independent decoupling and control of the Young's modulus (mechanical strength) and rheological behavior (viscoelasticity) of hydrogels can be successfully achieved. This demonstrates precise controllability of mechanical properties and successful decoupling of mechanical strength and viscoelastic characteristics.

[0161] Example 4 Preparation and characterization of heparin-loaded dual-network microspheres (Hep-MS) (1) Heparin grafting Microspheres prepared according to the method in Example 2 (containing 1.6% (w / v) GelMA (95% substitution degree) and 2.3% (w / v) GelMA (15% substitution degree)) were dispersed in MES buffer (pH 5.5) containing 5 mg / mL heparin sodium. EDC (2 mg / mL) and NHS (2 mg / mL) were added, and the mixture was activated for 30 minutes. The pH was adjusted to 7.4, and the reaction was carried out at room temperature for 12 hours. The microspheres were then washed to obtain Hep-dECM / GelMA microspheres.

[0162] (2) Grafting rate detection (toluidine blue method) The grafting amount was determined using the principle of specific binding between toluidine blue (TBO) and heparin. This embodiment employs the toluidine blue (TBO) colorimetric method to quantitatively determine the heparin loading on the surface of microspheres. A quantitative amount of microspheres were immersed in toluidine blue staining solution (0.4 mg / mL TBO, 2 mg / mL NaCl, 0.1 M HCl) and incubated at room temperature for 2 hours to allow the dye to specifically complex with heparin molecules. The staining solution was removed, and the microspheres were repeatedly washed with ultrapure water until the supernatant was clear and colorless to remove unbound dye. Eluent (ethanol:0.1 M NaOH = 4:1, v / v) was added, and the precipitated dye complex was dissolved by shaking. The absorbance of the supernatant was measured at 530 nm using a microplate reader, and the heparin grafting amount per unit mass of microspheres was calculated according to a preset standard curve.

[0163] The results showed that the heparin grafting density of Hep-MS microspheres was approximately 1.8 ± 0.1 μg / mg. Figure 6 The control group without grafting showed almost no TBO adsorption, confirming the effectiveness of chemical grafting.

[0164] Example 5 Biocompatibility Detection of Dual-Network Microspheres (Hep-MS) (1) Cell proliferation detection Microspheres from Example 4, both unloaded (MS) and heparin-loaded (Hep-MS) microspheres, which had been fully swollen in PBS, were extracted at 0.2 mg / ml in cell culture medium (DMEM + 10% fetal bovine serum) at 37°C for 24 hours to obtain the extract. The control group used normal cell culture medium (DMEM + 10% fetal bovine serum). The microsphere extract was seeded into 96-well plates (experimental group), and mouse hippocampal neuron (HT-22) cell suspensions were inoculated and cultured for 1 and 3 days, respectively. At the detection time point, the culture medium was aspirated, and fresh culture medium containing 10% (v / v) Cell Counting Kit-8 (CCK-8) reagent (100 μL per well) was added. After incubation at 37°C in the dark for 1 hour, the absorbance (OD value) at 450 nm was measured using a microplate reader to reflect cell proliferation activity. Results are as follows: Figure 7 As shown, the results indicated that after 3 days of culture, the cell viability of the Hep-MS group was significantly improved, reaching approximately 145% of that of the control group, demonstrating that the microspheres have excellent biocompatibility and the ability to promote cell proliferation.

[0165] (2) Observation of cytoskeleton morphology (immunofluorescence staining) Mouse hippocampal neurons (HT-22) were cultured for 2 days on the surface of dual-network hydrogel microspheres (DMEM medium + 10% fetal bovine serum, 37°C 5% CO2 cell culture incubator).

[0166] After washing cells with PBS, they were fixed with 4% paraformaldehyde (PFA) at room temperature for 10-30 minutes; after washing with PBS, they were permeabilized with 0.5% Triton X-100 solution for 5 minutes. The cells were washed three times with PBS (5 minutes each time). FITC-labeled phalloidin working solution was added and incubated at room temperature in the dark for 30 minutes to label F-actin. After washing with PBS, the cell nuclei were counterstained with DAPI solution. Finally, after washing with PBS, the green fluorescent cytoskeleton and blue fluorescent cell nuclei were observed under a fluorescence microscope through corresponding filters to assess the cell spreading morphology on the microsphere surface. Results are as follows: Figure 8As shown, the cells spread well on the surface of the microspheres and are accompanied by significant neural synaptic extension.

[0167] Example 6 In vitro degradation experiment The dual-network hydrogel of Example 3 was weighed (wet weight approximately 1g) and placed in PBS buffer, and incubated on a shaker at 37°C.

[0168] On day 7, a sample from formulation B was taken out, washed with deionized water, freeze-dried, and its microstructure changes were observed using SEM. The SEM image is shown below. Figure 9 As shown, the second network degrades first.

[0169] The retention mass of each group of dual-network hydrogels after degradation over time is as follows: Figure 10 As shown, in the dual-network hydrogel of the present invention, the rapid degradation of the second network within 7-14 days corresponds to the window period of acute inflammation subsiding and microvascular formation at the site of injury; subsequently, the slow degradation of the first network for 30-60 days is to address the characteristics of slow growth of central nerve axons and delayed secretion of autologous matrix; to prevent the lesion site from being filled with dense gel scars, and to ensure the construction of a continuous and stable regenerative microenvironment before the newly generated nerve tissue completely fills the defect.

[0170] In this invention, the photocurable second network provides initial mechanical support, ensuring the structural integrity of the microcarrier during the initial injection and implantation stages. Subsequently, this network preferentially degrades to reduce physical steric hindrance and promote deep cell infiltration, while the slower-degrading dECM first network serves as a long-lasting biological scaffold to continuously induce neural regeneration. The second network is designed to preferentially degrade in vivo, a process that releases micron-sized pores in situ within the microcarrier, significantly reducing physical steric hindrance and facilitating the deep infiltration of host cells and new blood vessels into the microspheres. The remaining dECM network, with its slower degradation rate, can serve as a long-lasting bioactive template, continuously providing neural induction signals, thereby achieving a dynamic match between the material degradation rate and the neural tissue regeneration cycle.

[0171] Furthermore, the surface-coated heparin utilizes specific affinity to inhibit the burst release effect of growth factors. The binding of heparin to growth factors stabilizes the tertiary structure of these factors, preventing their rapid enzymatic inactivation in vivo. This microcarrier can maintain an effective drug concentration for more than 14 days, covering the critical window period of nerve repair, thereby achieving long-term sustained drug release and dynamically matching the material degradation rate, changes in mechanical properties, and the tissue repair cycle.

[0172] Comparative Example 1 The formulation and preparation method are basically the same as those in Example 3 Group B, except that the second network photocrosslinking step is performed first (405 nm blue light (light intensity 30 mW / cm²)). 2After irradiation for 40 seconds, the first network thermal crosslinking step (heated to 37°C and held for 10 min) is performed.

[0173] Scanning electron microscope images of the obtained materials are as follows Figure 11 As shown, the average pore size is 36.52 ± 8.43 μm. It can be seen that after changing the crosslinking sequence, the formation of the first network of the cell matrix could not be observed, indicating that the formation of the second network first hinders the thermal assembly of the first network. The following is a scanning electron microscope image of the fiber network after gelation of only the first network of the decellularized matrix: Figure 12 As shown, the first network has a mesh-like structure, while the second network mainly has a sheet-like structure.

[0174] Comparative Example 2 Group A. Spinal cord chemoenzymatic hydrolysis dECM prepared using the same method as in Example 1; Group B. Spinal cord dECM enzymatically hydrolyzed lyophilized powder prepared by the same method as in Example 1.

[0175] The chemically hydrolyzed dECM from group A and the lyophilized dECM from group B were dissolved in 0.01 M HCl solution at 1% (w / v) for 2 hours until completely dissolved. The pH was then adjusted to 7.4 with NaOH containing phenol red acid-base indicator and allowed to stand at 37°C for 20 minutes.

[0176] Experimental results are as follows Figure 13 As shown, the right side represents group A (without pepsin hydrolysis), and the left side represents group B (with pepsin hydrolysis). It can be seen that the spinal cord decellularized matrix without further pepsin hydrolysis cannot undergo temperature-induced self-assembly to form a gel, while the spinal cord decellularized matrix of this invention, after further pepsin hydrolysis, can successfully form a gel. A possible reason is that in the spinal cord decellularized matrix obtained directly through chemical hydrolysis, large protein molecules such as collagen are bound within the decellularized matrix particles due to hydrogen bonds and physical entanglement. Simply relying on the disruption of hydrogen bonds under acidic conditions is insufficient to completely dissolve them, preventing effective release and dispersion, and thus hindering subsequent controllable self-assembly to form a gel. In contrast, pepsin, through limited enzymatic hydrolysis, degrades physically entangled and difficult-to-disperse large protein molecules into easily soluble and dispersed peptides with slightly smaller molecular weights, weakening intermolecular interactions under acidic conditions and exposing hydrophilic groups. This significantly improves the dispersibility and solubility of protein molecules, allowing them to self-assemble into a gel under temperature-induced conditions through hydrophilic-phobic interactions and hydrogen bonding.

[0177] Comparative Example 3 Group A: 5% (w / v) GelMA (30% degree of substitution), dissolved in phosphate-buffered saline (PBS) containing 0.25% (w / v) photoinitiator (LAP), was directly used as an aqueous precursor solution and injected into the mold. 405 nm blue light (30 mW / cm² intensity) was then applied. 2 Irradiate for 40 seconds to cure the photogel into a large cylindrical hydrogel (20mm in diameter and 1.5mm in height).

[0178] Group B: 5% (w / v) GelMA (30% degree of substitution) + 1% (w / v) dECM enzymatically hydrolyzed lyophilized powder, hydrogel blocks (20 mm in diameter, 1.5 mm in height) were prepared according to the same method as in Example 3: Compared with group A (GelMA), which showed an increasing Young's modulus (850 Pa) and loss modulus (19.12 Pa (1 Hz)), group B (GelMA+dECM) exhibited both an increasing Young's modulus (1120 Pa) and loss modulus (27.3 Pa (1 Hz)). This indicates that the first network primarily enhances the overall skeletal strength through physical entanglement. More importantly, this non-covalent physical cross-linking increases the loss modulus of the gel, improves the dynamic viscoelastic properties of the hydrogel, and endows the matrix with superior dynamic remodeling capabilities. This allows the gel to dissipate and yield when facing the traction forces of nerve cell growth and neurite extension, significantly reducing the physical resistance to cell inward growth, thus making it more conducive to cell spreading and matrix remodeling.

[0179] discuss The inventors' research indicates that the physicomechanical properties of the extracellular microenvironment, particularly the Young's modulus and viscoelasticity of the matrix, are key factors influencing cell behavior. Through mechanotransduction mechanisms, the mechanical signals of the matrix can directly regulate the adhesion, migration, and differentiation direction of stem cells. For example, neural tissue typically requires a low Young's modulus to facilitate neuronal growth, while also needing specific viscoelasticity to support axonal extension. However, the mechanical properties of existing single dECM carriers are often fixed and unadjustable, and their strength is insufficient to meet load-bearing / shear resistance requirements. While chemical cross-linking can improve strength, viscoelasticity is lost, the material hardens, loses its adhesion and flexibility to tissues, and may even affect cell adhesion and differentiation, making it difficult to accurately match the differentiated mechanical support requirements of different repair stages.

[0180] In addition, biodegradability is an important feature for achieving functional repair. The degradation rate needs to match the regeneration rate of the host tissue to avoid scaffold residue affecting regeneration or scaffold degradation leading to tissue collapse.

[0181] Furthermore, to promote the repair of damaged tissues, it is often necessary to load bioactive molecules (such as neurotrophic factors) onto a carrier. Conventional techniques mostly employ physical embedding or physical adsorption for loading. Due to the lack of a specific binding mechanism, growth factors tend to diffuse rapidly through the carrier pores, leading to a "burst release effect" in the early stages of application, making it difficult to meet the requirements of long-term, sustained release of bioactive molecules for processes such as nerve regeneration.

[0182] In summary, developing a self-sourced active microcarrier that can retain the rich biological activity of dECM, possess tunable mechanical strength, degradation characteristics, and biocompatibility to meet the needs of different tissues, and achieve long-term delivery of bioactive molecules is a technical problem that needs to be solved in the field of biomaterials and is of great significance.

[0183] It should be understood that after reading the above teachings of this invention, those skilled in the art can make various alterations or modifications to this invention, and these equivalent forms also fall within the scope defined by the appended claims.

Claims

1. A method for preparing a decellularized matrix dual-network hydrogel, characterized in that, Includes the following steps: S1. Preparation of aqueous precursor solution: Dissolve the decellularized matrix in water under acidic conditions, then mix it with an aqueous solution of (meth)acrylamide gelatin and photoinitiator, and adjust the pH to neutral at a low temperature of 0-10℃ to obtain an aqueous precursor solution; S2. Primary thermal gelation: The aqueous precursor solution is heated to 35-40°C to induce primary thermal gelation of the decellularized matrix, forming a mixture containing a first network gel; and S3. Photocrosslinking and curing: Applying light of a predetermined wavelength to the mixture containing the first network gel initiates a polymerization reaction of (meth)acrylamide gelatin to form a second network, thereby obtaining an interpenetrating double network hydrogel; The decellularized matrix is ​​prepared by the following method: Step i: Cell membrane disruption: Fresh spinal cord tissue was cut into pieces and frozen at -60°C to -80°C, and repeated freeze-thaw cycles were performed to disrupt the cell membrane. Step ii: Chemically enzymatically hydrolyze the product from step i to obtain the chemically enzymatically hydrolyzed product; Step iii: Add the chemical enzymatic hydrolysis product obtained in step ii to a 0.008-0.015 mol / L HCl solution containing 0.3-1 mg / mL pepsin at a mass fraction of 0.1-1% w / w. Stir continuously at 20-38°C until completely dissolved. Freeze-dry the pepsin hydrolysis product to obtain the decellularized matrix. The chemical enzymatic hydrolysis includes the following steps: repeating the following treatment steps on the tissue obtained in step i for 2-4 cycles: Water treatment for 4-8 hours, followed by treatment with 2-4 v / v% Triton X-100 solution for 5-8 hours; Water treatment for 15-30 min × 3 times, followed by treatment with 0.03-0.05 g / mL sodium deoxycholate solution for 8-12 h; and Water treatment for 15-30 minutes × 3 times, followed by DNase treatment for 3-5 hours; The chemical enzymatic hydrolysis product is then obtained by freeze-drying after the following cleaning steps: Water treatment for 15-30 min × 3 times, 3-5 v / v% ethanol treatment for 3-5 h, and water treatment for 15-30 min × 3 times; All processing steps were carried out in a constant temperature shaker at 37±2℃ with a rotation speed of 100-200 rpm. After each reagent was processed, the liquid was discarded before proceeding to the next operation. In the aqueous precursor solution, the concentration of the decellularized matrix is ​​7.5-20 mg / mL; The concentration of the (meth)acrylamide gelatin is 1-20 w / v%; and the (meth)acrylamide gelatin is one or more mixtures with a degree of substitution selected from 5-95%. The average pore size of the dual-network hydrogel is 10-30 μm, and the Young's modulus is 0.5 kPa-2.5 kPa.

2. The preparation method according to claim 1, characterized in that, The method further includes: S4. Functional modification: covalently grafting heparin molecules onto the amino groups on the surface of the dual-network hydrogel via an amidation reaction.

3. The preparation method according to claim 1, characterized in that, In step S1, the (meth)acrylamide gelatin is selected from methacrylamide gelatin and acrylamide gelatin.

4. The preparation method according to claim 1, characterized in that, The method includes the following steps: S1. Preparation of aqueous precursor solution: Dissolve the decellularized matrix in water under acidic conditions, then mix it with an aqueous solution of (meth)acrylamide gelatin and photoinitiator, adjust the pH to neutral at a low temperature of 0-10℃ to obtain an aqueous precursor solution, and store it at a low temperature of 0-10℃. S2. Primary thermogelation: Using the aqueous precursor solution as the dispersed phase and an oily liquid containing a surfactant as the continuous phase, droplet generation is performed in a microfluidic chip. Primary thermogelation of the decellularized matrix is ​​induced at the rear end of the chip by heating to 35-40℃, forming microspheres containing a first network gel. S3. Photocrosslinking and curing: Apply light of a predetermined wavelength to the mixture microspheres containing the first network gel to initiate the polymerization reaction of (meth)acrylamide gelatin to form a second network, thereby obtaining interpenetrating double-network hydrogel microspheres.

5. The preparation method according to claim 2, characterized in that, In step S4, the amidation reaction involves activating the carboxyl group in the heparin molecule with EDC / NHS and then reacting it with the amino group on the surface of the dual-network hydrogel.

6. The preparation method according to claim 1, characterized in that, Step S1 includes the step of: adjusting the mechanical properties of the dual-network hydrogel by adjusting the concentration and / or degree of substitution of the (meth)acrylamide gelatin in the aqueous precursor solution.

7. The preparation method according to claim 1, characterized in that, The decellularized matrix is ​​derived from humans, rats, mice, monkeys, or pigs.

8. A decellularized matrix dual-network hydrogel, characterized in that, The decellularized matrix dual-network hydrogel is prepared by the preparation method according to any one of claims 1-7.

9. A cell culture substrate or tissue engineering scaffold, characterized in that, It comprises: the decellularized matrix dual-network hydrogel as described in claim 8; wherein the decellularized matrix dual-network hydrogel has a covalently modified heparin layer and is loaded with cytokines having heparin binding domains.

10. The use of the decellularized matrix dual-network hydrogel as described in claim 8 in the preparation of tissue engineering scaffolds, drug carriers and / or cell culture substrates.

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