Rigidity-adjustable injectable hydrogel as well as preparation method and application thereof

Abstract

The invention discloses injectable hydrogel with adjustable rigidity as well as a preparation method and application of the injectable hydrogel. A precursor solution of the hydrogel comprises the following components in percentage by mass: 0-0.1% (not equal to 0) of carboxylated multi-walled carbon nanotubes, 5-15% of gelatin methacrylate, 0.1-1% of a photoinitiator and the balance of a PBS (Phosphate Buffer Solution); the rigidity of the hydrogel can be regulated and controlled through the content of the carboxylated multi-walled carbon nanotubes. Due to the excellent injectability, flowability, biocompatibility and photosensitive responsiveness of the hydrogel, the hydrogel can be subjected to in-situ photo-crosslinking after being injected in vivo, can be formed in situ, accurately adapts to and fills complex and irregular tissue defect areas, and has good mechanical properties and tissue repair effects. Besides, the hydrogel can simulate a tissue cell microenvironment and establish bionic systems with different rigidities, provides possibility for directional differentiation of MSCs, is particularly suitable for precise repair of complex bone defects, and has important application value in the field of bone tissue engineering.

Description

Technical Field

[0001] The present invention relates to the technical field of tissue engineering materials, and in particular to an injectable hydrogel with adjustable stiffness, a preparation method thereof and an application thereof. Background Art

[0002] As an important scaffold of the human body, the skeletal system not only supports and protects other tissues and organs, but also participates in mineral metabolism and hematopoiesis. The regeneration of severe bone defects caused by trauma, tumors or infection remains a major challenge in the field of orthopedics. In addition, the bone defect area often has an irregular shape, and traditional bone replacement materials are difficult to match it precisely. Although some injectable bone cements have good fluidity and biocompatibility and can adapt to complex defect structures, they lack osteoinductivity and are non-degradable. In recent years, tissue engineering technology has developed rapidly. By simulating the extracellular matrix microenvironment to promote cell adhesion, proliferation and differentiation, it has provided new ideas for the repair of bone defects and is expected to achieve more efficient bone regeneration.

[0003] Tissue engineering, an emerging branch of regenerative medicine, has recently shown promising applications in reconstructing damaged tissues. This technology uses cells as its fundamental building blocks, combined with scaffold materials to provide structural support, guide cell growth and matrix synthesis, and achieve functional tissue regeneration through biomimetic strategies that mimic the physiological environment of native tissue. Numerous studies have demonstrated that when cells are placed or recruited into a scaffold, they sense a variety of mechanical cues, including the material's stiffness, contractility, surface microstructure, and spatial dimensions. These cues often interact with other physicochemical and biochemical factors through receptors on the cell membrane. This complex network of extracellular signals activates a series of intracellular signaling pathways, ultimately translating into changes in cellular behavior, such as migration, proliferation, and differentiation. Furthermore, during tissue regeneration, adaptation, and disease progression, cells exert forces on the scaffold, reshaping its microstructure. Scaffold material stiffness is considered a key mechanical property influencing cell fate. Over the past few decades, substrate stiffness has been extensively studied and is recognized as a key factor in regulating cell behavior and inducing stem cell lineage differentiation. Matrix stiffness plays a key role in stem cell differentiation. Different stiffnesses simulate different tissue environments and determine the direction of stem cell differentiation (Engler, Adam J. et al. Matrix Elasticity Directs Stem Cell Lineage Specification. Cell, 2006, Volume 126, Issue 4, 677-689). Mesenchymal stem cells (MSCs) were cultured on hydrogels of different stiffnesses, and it was found that MSCs responded significantly to matrix stiffness. On a soft matrix that simulates brain elasticity (0.1-1 kPa), MSCs were branched, similar to the morphology of neurons; on a matrix that simulates muscle elasticity (8-17 kPa), MSCs were spindle-shaped, similar to myoblasts; on a matrix that simulates bone elasticity (25-40 kPa), MSCs were polygonal, close to the morphology of osteoblasts. This shows that different matrix stiffnesses promote the development of MSCs into different lineage differentiation morphologies. Therefore, if we can simulate the cell microenvironment and flexibly regulate the stiffness of the matrix, we can achieve the regulation of directional pre-differentiation of MSCs in vivo, thereby repairing specific tissues and further promoting the application of MSCs in the field of regenerative medicine.

[0004] Gelatin molecular chains are rich in reactive groups such as amino and hydroxyl groups. When polymerizable methacryloyl groups are introduced, gelatin methacryloyl (GelMA) is formed. The introduction of methacryloyl groups allows GelMA to undergo free radical polymerization under ultraviolet light with the aid of a photoinitiator, forming a hydrogel with a three-dimensional cross-linked network structure. In recent years, this hydrogel has attracted considerable attention in the fields of tissue engineering and regenerative medicine. However, similar to traditional hydrogel systems, pure GelMA hydrogels generally suffer from poor mechanical properties and fragility, limiting their application in tissue regeneration under high loads. Currently, many researchers have modified the properties of GelMA hydrogels. For example, invention patent CN116077736A discloses an injectable bone repair scaffold based on porous gelatin composite microspheres and its preparation method. Methacrylic acid-modified gelatin and inorganic nanoparticles are combined to form an organic / inorganic solution. Polyethylene oxide is added to the organic / inorganic solution, mixed evenly, and then emulsified to produce microspheres. However, inorganic nanoparticles and GelMA are mainly combined through physical mixing, with weak interfacial interactions and limited mechanical enhancement. HA particles tend to aggregate in GelMA, leading to uneven stress distribution, and high HA content may destroy the network structure of GelMA, which in turn reduces mechanical properties. Therefore, the ability to improve the mechanical properties of GelMA through inorganic nanoparticles and to regulate the stiffness of stem cell differentiation through the content of inorganic nanoparticles is limited. At the same time, there is a lack of coordinated design for cell adhesion, signal transduction, and microenvironmental biomimetic, making it difficult to meet the multiple demands for mechanical support, directional cell differentiation, and tissue regeneration in the repair of complex bone defects. Summary of the Invention

[0005] In view of the above-mentioned deficiencies in the prior art, the technical problem to be solved by the present invention is: how to provide an injectable hydrogel with excellent mechanical properties, good injectability and biocompatibility, and continuously adjustable stiffness characteristics, as well as its preparation method and application. The hydrogel should be able to effectively make up for the problems of insufficient mechanical properties of the traditional GelMA system, weak interface bonding of inorganic particles and limited stiffness control range, and achieve precise guidance of directional differentiation of MSCs on the basis of simulating different tissue microenvironments, thereby meeting the needs of different tissue engineering repairs.

[0006] To address the above technical issues, the present invention adopts the following technical solution: The present invention provides an injectable hydrogel with adjustable stiffness. The hydrogel is a network structure formed by photocrosslinking a precursor solution. The precursor solution of the hydrogel comprises the following components by weight: 0-0.1% (not 0) carboxylated multi-walled carbon nanotubes, 5-15% methacrylated gelatin, 0.1-1% photoinitiator, and the balance PBS solution. The stiffness of the hydrogel can be controlled by adjusting the content of the carboxylated multi-walled carbon nanotubes. This composition ratio ensures successful photocrosslinking while maintaining the fluidity of the hydrogel upon injection, enabling the repair of irregular tissue defects.

[0007] GelMA has a cell adhesion site arginine-glycine-aspartic acid (RGD), which is a tripeptide that facilitates certain cell activities such as adhesion, migration and differentiation into various lineages, as well as a matrix metalloproteinase (MMP) sequence, which is an endopeptidase that supports enzyme degradation and plays a key role in skin wound healing, morphogenesis and tissue repair. After carboxylation, CNTs can form a stable connection with methacrylate gelatin at the molecular level through hydrogen bonds, electrostatic interactions and even covalent bonds, which can effectively improve the mechanical strength and network uniformity. In addition, due to the presence of long carbon chains in carboxylated multi-walled carbon nanotubes and the ubiquitous sp 2 Hybridization, CNTs act as a reinforcing agent in the GelMA network. The mechanical properties of GelMA hydrogels can be enhanced by functionalizing carbon nanotubes with multi-walled -COOH groups, providing a fiber structure in the interconnected sponge-like hydrogel network in the presence of carbon nanotubes. Due to the formation of a gel-like nanofiber mesh network coated with multi-walled carbon nanotubes, the modulus of the hydrogel is significantly improved. Moreover, the nanofiber units formed after adding CNTs can also retain the beneficial bioactive properties of GelMA, such as bioactivity, highly porous morphology and biodegradability. In addition, different amounts of CNTs are doped in the gel to adjust its stiffness to meet the needs of different tissue repairs, promote MSC differentiation into the desired tissue, adsorb growth factors such as BMP and affect the immune microenvironment (such as promoting M2 macrophage polarization). It is more systematic at the level of tissue regeneration regulation, including a stiffness range that promotes osteogenic differentiation of stem cells, and the addition of trace amounts of CNTs will not be toxic to cells.

[0008] Preferably, the carboxylated multi-walled carbon nanotubes have a diameter of 8 to 15 nm and a length of 45 to 55 μm.

[0009] Preferably, the photoinitiator is Irgacure 2925 or LAP.

[0010] Another object of the present invention is to provide a method for preparing the above-mentioned injectable hydrogel with adjustable stiffness, comprising the following steps:

[0011] 1) introducing carbon nanotubes into a strong acid solution and subjecting them to water bath sonication to obtain a mixed solution, wherein the strong acid is a mixture of concentrated sulfuric acid and concentrated nitric acid; thus, the surface of the carbon nanotubes is carboxylated by the strong acid solution to improve their dispersibility in aqueous solution. The oxidizing power of the mixed acid is stronger than that of a single acid. If only a single acid is used, the carbon atoms cannot be oxidized, and no carboxylated carbon nanotubes can be obtained.

[0012] 2) diluting the mixed solution obtained in step 1) with water, then continuing to add alkaline solution to adjust the pH value of the mixed solution to 7.0, then filtering and vacuum drying the mixed solution to a constant weight to obtain surface carboxylated carbon nanotubes;

[0013] 3) dispersing the carboxylated carbon nanotubes obtained in step 2) in a PBS solution containing methacrylate gelatin, adding a photoinitiator, uniformly mixing the obtained precursor solution, and then photocrosslinking to form the injectable hydrogel.

[0014] Preferably, the volume ratio of concentrated sulfuric acid to concentrated nitric acid in the strong acid solution in step 1) is 3:1.

[0015] Preferably, the temperature of the water bath ultrasound in step 1) is 60-80° C., and the time is 4-6 hours.

[0016] Another object of the present invention is to provide the use of the above-mentioned injectable hydrogel with adjustable stiffness or the injectable hydrogel with adjustable stiffness prepared by the above-mentioned method in the preparation of tissue regeneration and repair materials, including but not limited to bone tissue regeneration and repair materials.

[0017] Another object of the present invention is to provide a hydrogel scaffold for tissue repair. The hydrogel scaffold can be obtained by injecting a precursor solution of the above-mentioned injectable hydrogel with adjustable stiffness into a mold or the location of a tissue defect, and then irradiating it with ultraviolet light or visible light and then photocrosslinking it in situ. In this way, the precursor solution of the hydrogel is filled into the tissue defect site by injection, maintaining fluidity during the injection process, and the gelatin with methacryloyl groups can undergo free radical polymerization under the action of a photoinitiator. Under ultraviolet or visible light irradiation, the photoinitiator generates free radicals, which trigger the polymerization of methacryloyl groups, forming a three-dimensional network structure after in situ photocrosslinking. In addition, this injectable gel system can regulate stiffness to promote the differentiation of mesenchymal stem cells into different tissues, thereby achieving the repair of irregular tissues.

[0018] Preferably, the wavelength of the ultraviolet light is 300-400 nm, the power is 15-50 W, and the irradiation time is 1-10 min.

[0019] Compared with the prior art, the present invention has the following beneficial effects:

[0020] 1. The present invention provides an injectable hydrogel with adjustable stiffness, which uses GelMA as the base material for repair materials and dopes the gel with different contents of carboxylated multi-walled carbon nanotubes to adjust its stiffness. Carboxylated multi-walled carbon nanotubes can form stable connections with methacrylate gelatin at the molecular level through hydrogen bonds, electrostatic interactions, and even covalent bonds. On the one hand, CNTs provide a fiber structure in the sponge-like hydrogel network, forming a gel-like nanofiber mesh network coated with multi-walled carbon nanotubes, effectively improving the mechanical strength and network uniformity, and significantly improving the modulus of the hydrogel; on the other hand, the nanofiber units formed after adding CNTs can also retain the beneficial bioactive properties of GelMA, such as bioactivity, highly porous morphology, and biodegradability, which are beneficial to certain cell activities, such as adhesion, migration, and differentiation into various lineages, and play a key role in skin wound healing, morphogenesis, and tissue repair. The hydrogel of the present invention can flexibly regulate the stiffness of the hydrogel simply by adjusting the content of CNTs in the GelMA hydrogel. Its compression modulus can range from 12.09±4.01kPa to 49.42±10.99kPa, which can meet the needs of different tissue repair. By adjusting the content of carbon nanotubes in GelMA, the present invention can establish a biomimetic system with different stiffness, making it possible to direct the differentiation of MSCs and providing a new option for tissue repair.

[0021] 2. The injectable hydrogel with adjustable stiffness prepared by the present invention can mimic the cell microenvironment, including maintaining its effectiveness, self-renewal and differentiation, migration, proliferation and interaction with other cells, and builds a new platform for the study of matrix mechanics in bone tissue repair. It simulates the tissue cell microenvironment, establishes a bionic system with different stiffness, promotes the differentiation of MSCs, and the addition of CNTs does not affect the degradation performance, water absorption and porosity of GelMA. At the same time, the hydrogel also has sufficient mechanical properties and biological effects, and can adapt to the needs of different tissues by adjusting the stiffness. MSCs can be used to repair different tissues such as bone tissue, muscle tissue and adipose tissue, and the hydrogel system has no inhibitory or toxic effects on cells and their activity. It is a very potential tissue repair material. The present invention does not require complicated processes and operations, does not require expensive instruments, and the raw materials are simple and easy to obtain. The preparation cost is low, and it is convenient for large-scale preparation, with broad application prospects.

[0022] 3. The present invention also provides a hydrogel scaffold for tissue repair with adjustable stiffness, which is injectable and can be formed in situ. The hydrogel precursor solution can be injected to accurately fill irregular tissue defect areas, and has good fluidity and photosensitivity. After injection into the defect area, it can be rapidly cured in situ by ultraviolet or visible light to form a stable three-dimensional structure, ensuring the shape retention and mechanical support of the hydrogel at the implant site, and can induce in situ tissue regeneration. Before injection, the CNTs-GelMA system remains in a fluid state, making it easy to inject into complex or deep bone defects through a fine-diameter needle; after injection, it quickly solidifies and takes shape under irradiation conditions, forming a close fit with the surrounding tissue, thereby achieving seamless docking and in situ repair. In addition, the addition of CNTs enables the hydrogel to have better mechanical matching ability while maintaining injectability, realizing a composite functional system of injectability + stiffness regulation + in situ molding. This combination strategy has significant advantages in adapting to complex tissue defects and inducing stem cell differentiation. It can be adjusted according to the stiffness requirements of the target tissue and simulate different tissue microenvironments. It is particularly suitable for the growth and differentiation of osteoblasts in bone tissue that are sensitive to stiffness. BRIEF DESCRIPTION OF THE DRAWINGS

[0023] Figure 1 GelMA prepared in the present invention; (A) Schematic diagram of the chemical synthesis of GelMA; (B) Actual picture of GelMA.

[0024] Figure 2 Schematic diagram of CNTs-GelMA hydrogel scaffolds with different stiffness prepared in the present invention; (A) hydrogel precursor solution; (B) gelation of the precursor solution; (C) actual image of the CNTs-GelMA scaffold.

[0025] Figure 3 The stiffness of CNTs-GelMA hydrogel scaffolds with different stiffness prepared in the present invention; (A) stress-strain curve; (B) compression modulus.

[0026] Figure 4 The microstructure of CNTs-GelMA hydrogel scaffolds with different stiffness prepared in the present invention.

[0027] Figure 5 The hydrophilic properties of CNTs-GelMA hydrogels with different stiffness prepared in the present invention; (A) Swelling rate of CNTs-GelMA scaffolds with different stiffness; (B) Water absorption of CNTs-GelMA scaffolds with different stiffness; (C) Water retention rate of CNTs-GelMA scaffolds with different stiffness.

[0028] Figure 6 Live / dead staining of cells seeded on CNTs-GelMA hydrogel scaffolds with different stiffnesses.

[0029] Figure 7 Quantification of cell activity seeded on CNTs-GelMA hydrogel scaffolds with different stiffness. DETAILED DESCRIPTION

[0030] The present invention will be further described in detail below with reference to the following examples. Unless otherwise specified, the reagents and experimental methods used in the examples are commercially available and are conventional operations.

[0031] 1. Preparation method of an injectable hydrogel with adjustable stiffness and its tissue repair scaffold

[0032] Example 1: This example was prepared by the following steps:

[0033] 1) Selecting multi-walled carbon nanotubes with a diameter of 8 to 15 nm and a length of 45 to 55 μm, placing them in a strong acid solution obtained by mixing concentrated sulfuric acid and concentrated nitric acid in a volume ratio of 3:1, and then ultrasonically treating them in an 80°C water bath for 4 hours. After the reaction is completed, adding a large volume of deionized water to dilute the solution, and adjusting the pH value of the diluted solution to 7.0 with ammonia water to obtain a mixed solution.

[0034] 2) The mixed solution obtained in step 1) is filtered using a circulating water bath vacuum pump, and then vacuum dried to a constant weight to obtain surface carboxylated multi-walled carbon nanotubes.

[0035] 3) 10 g of gelatin was dissolved in 100 ml of phosphate buffered saline (PBS) and heated to 60° C., and then 8 ml of methacrylic anhydride was slowly added dropwise with vigorous stirring. After reacting for 3 h, 400 ml of PBS was added to stop the reaction, and the solution was dialyzed with deionized water in an 8000-14000 kDa dialysis bag at 40° C. for 7 days to remove any unreacted methacrylic anhydride. The solution was then placed in a -20° refrigerator overnight, then in a -80° refrigerator for one day, and finally freeze-dried in a freeze dryer for 3 days to produce fibrous white foam of gelatin methacrylate (GelMA), as shown in FIG. Figure 1 shown.

[0036] 4) The freeze-dried GelMA was dissolved in PBS for 30 min, and then surface carboxylated multi-walled carbon nanotubes were added and ultrasonically mixed in a water bath. Then, I2959 photoinitiator was added, and the GelMA and photoinitiator were completely dissolved to obtain a hydrogel precursor solution, wherein the content of carboxylated CNTs in the precursor solution was 0.05% (w / v), the content of methacrylated gelatin was 10% (w / v), and the content of the photoinitiator was 0.5% (w / v).

[0037] 5) The precursor solution was then transferred into a well plate and cured using a 365 nm UV lamp for 3 minutes to prepare a CNTs-GelMA scaffold.

[0038] Example 2: This example was prepared by the following steps:

[0039] 1) Selecting multi-walled carbon nanotubes with a diameter of 8 to 15 nm and a length of 45 to 55 μm, placing them in a strong acid solution obtained by mixing concentrated sulfuric acid and concentrated nitric acid in a volume ratio of 3:1, and then ultrasonically treating them in an 80°C water bath for 4 hours. After the reaction is completed, adding a large volume of deionized water to dilute the solution, and adjusting the pH value of the diluted solution to 7.0 with ammonia water to obtain a mixed solution.

[0040] 2) The mixed solution obtained in step 1) is filtered using a circulating water bath vacuum pump, and then vacuum dried to a constant weight to obtain surface carboxylated multi-walled carbon nanotubes.

[0041] 3) 10 g of gelatin was dissolved in 100 ml of phosphate-buffered saline (PBS) and heated to 60°C. 8 ml of methacrylic anhydride was then slowly added dropwise with vigorous stirring. After 3 h of reaction, 400 ml of PBS was added to stop the reaction. The solution was dialyzed against deionized water in an 8000-14000 kDa dialysis bag at 40°C for 7 days to remove any unreacted methacrylic anhydride. The solution was then placed in a -20°C refrigerator overnight, then in a -80°C refrigerator for one day, and finally lyophilized in a freezer for 3 days to produce fibrous white foam of gelatin methacrylate (GelMA).

[0042] 4) The freeze-dried GelMA was dissolved in PBS for 30 min, and then surface carboxylated multi-walled carbon nanotubes were added and ultrasonically mixed in a water bath. Then, I2959 photoinitiator was added, and the GelMA and photoinitiator were completely dissolved to obtain a hydrogel precursor solution, wherein the content of carboxylated CNTs in the precursor solution was 0.1% (w / v), the content of methacrylated gelatin was 10% (w / v), and the content of the photoinitiator was 0.5% (w / v).

[0043] 5) The precursor solution was then transferred into a well plate and irradiated with a 365 nm UV lamp for 3 min to gelate the solution, thereby preparing a CNTs-GelMA scaffold.

[0044] Comparative Example 1: In this comparative example, no carboxylated multi-walled carbon nanotubes were added, and other steps were the same as in Example 1.

[0045] 2. Performance Verification

[0046] 1. Observe the morphology of the hydrogels and scaffolds prepared in Examples 1-2 and Comparative Example 1. Figure 2 shown.

[0047] from Figure 2It can be seen that the precursor solutions of the hydrogels of Example 1 and Example 2 are ink-colored, and the carboxylated multi-walled carbon nanotubes are evenly distributed in the precursor solutions of the hydrogels ( Figure 2 A). The precursor solution in the centrifuge tube was placed under a 365nm UV lamp for 3 minutes to gel. It can be seen that both Examples 1 to 2 and Comparative Example 1 can be effectively cross-linked into a gel under the action of the photoinitiator ( Figure 2 B). Moreover, the CNTs-GelMA scaffold structure prepared in Examples 1 and 2 is more stable ( Figure 2 C), this is because CNTs have a high specific surface area and good mechanical properties. They can interact with the GelMA molecular chains, making the formed hydrogel scaffold more stable. In addition, during the photocrosslinking process, CNTs can serve as physical crosslinking points to enhance the network structure formed by GelMA, thereby increasing the strength and stability of the gel.

[0048] 2. The CNTs-GelMA scaffolds prepared in Examples 1 and 2 and Comparative Example 1 were subjected to mechanical compression performance tests. The results are as follows: Figure 3 shown.

[0049] from Figure 3 As can be seen in Figure A, under the same strain conditions, the stress of the scaffolds of Comparative Example 1, Example 1, and Example 2 varies in a gradient, with Comparative Example 1 having the lowest stress. This indicates that the addition of CNTs significantly improves the mechanical properties of GelMA hydrogel, and that the scaffold stress increases with increasing CNT concentration.

[0050] Furthermore, the compression modulus of the scaffolds of comparative example 1 and examples 1-2 at 20-30% strain was compared. Figure 3 B. The results show that the compression moduli of the CNTs-GelMA scaffolds of Comparative Example 1, Example 1, and Example 2 were 12.09±4.01kPa, 29.78±2.51kPa, and 49.42±10.99kPa, respectively. This indicates that the stiffness of the GelMA hydrogel can be successfully regulated by changing the CNTs concentration, with the regulation range being between 12.09 and 49.42kPa. Among them, a stiffness range of 25 to 40kPa can promote the differentiation of stem cells into osteoblasts.

[0051] 3. The CNTs-GelMA scaffolds of Comparative Example 1 and Examples 1-2 were analyzed by scanning electron microscopy to observe their microstructures. The results are as follows: Figure 4 shown.

[0052] As can be seen from the figure, Comparative Example 1 and Examples 1-2 have similar microstructures and pore distributions, indicating that the doping of a small amount of CNTs will not affect the overall microstructure of GelMA.

[0053] 4. The water absorption and water retention performance of the CNTs-GelMA scaffolds of Comparative Example 1 and Examples 1-2 were tested. The results are as follows: Figure 5 shown.

[0054] from Figure 5 As can be seen from A, after sufficient water absorption, the swelling rates of the CNTs-GelMA scaffolds in Comparative Example 1, Example 1, and Example 2 were 510.94±7.03%, 500.01±9.91%, and 484.93±0.2%, respectively. Compared to Comparative Example 1, there was no significant difference in the swelling rate in Example 1, while the swelling rate in Example 2 decreased slightly by only 5%. This may be due to altered interactions between CNTs and GelMA, affecting the movement and spatial arrangement of GelMA molecular chains and hindering water absorption. Alternatively, the higher CNT concentration may alter the hydrogel's pore structure, hindering water ingress and diffusion, but this effect is minor.

[0055] from Figure 5 As can be seen from Figure B, the average water absorption rates of the CNTs-GelMA scaffolds of Comparative Example 1, Example 1 and Example 2 are (6.47±0.34) g / g, (6.41±0.21) g / g and (6.62±0.26) g / g, respectively, all of which have a certain water absorption capacity. Figure 5 As shown in Figure C, the average water retention rates of the CNTs-GelMA scaffolds in Comparative Example 1, Example 1, and Example 2 were (1.028±0.0033) g / g, (1.018±0.0025) g / g, and (1.019±0.0011) g / g, respectively, indicating that all have a certain water retention capacity. This indicates that the addition of CNTs does not affect the water absorption and water retention of GelMA.

[0056] 5. The hydrogel scaffolds prepared in Comparative Example 1 and Examples 1-2 were subjected to cytotoxicity experiments. L929 cells were seeded on the hydrogel scaffolds of Comparative Example 1, Example 1, and Example 2, respectively. After 1, 3, and 7 days of seeding, the L929 cells on the hydrogel scaffolds were subjected to Live / Dead staining, and the cell staining was observed under a microscope. The results are shown in FIG. Figure 6 .

[0057] from Figure 6As can be seen, at the same implantation time, the green fluorescence (live cell) area accounts for a large proportion in both Comparative Example 1 and Examples 1-2, while the number of red fluorescence (dead cells) is small or absent, with no significant difference. Furthermore, as the implantation time increases, the green fluorescence in Comparative Example 1 increases, indicating an increase in the proportion of live cells, i.e., a good natural cell growth state. Simultaneously, the proportion of live cells in Examples 1 and 2 also increases, indicating that the cells in Comparative Example 1, Examples 1, and 2 have good proliferation capabilities. Therefore, the hydrogel systems in Examples 1 and 2 do not produce significant inhibitory or toxic effects on cells.

[0058] Furthermore, MTT was used to test the activity of L929 cells grown on the scaffolds of Comparative Example 1, Example 1 and Example 2 for 1 to 3 days, and the cell group cultured with only complete medium was used as a blank control, and its cell activity was set as 100%. The results are as follows: Figure 7 shown.

[0059] The results showed that after one day of seeding on the hydrogel scaffolds of Comparative Example 1, Example 1, and Example 2, the L929 cell viability was 69.8±1.67%, 89.11±0.7%, and 120.09±0.55%, respectively. The cell viability on the hydrogel scaffolds of Example 1 and Example 2 was significantly higher than that of Comparative Example 1. After two days of seeding on the hydrogel scaffolds of Comparative Example 1, Example 1, and Example 2, the L929 cell viability was 74.08±1.07%, 98.21±1.85%, and 95.27±4.92%, respectively. After three days of seeding on the hydrogel scaffolds of Comparative Example 1, Example 1, and Example 2, the L929 cell viability was 89.24±2.04%, 91.38±3.34%, and 80.81±1.91%, respectively. This indicates that the addition of a small amount of CNTs has no effect on cell viability.

[0060] The above description is only a preferred embodiment of the present invention and does not limit the present invention. Any modifications, equivalent substitutions and improvements made within the spirit and principles of the present invention should be included in the scope of protection of the present invention.

Claims

1. An injectable hydrogel with adjustable stiffness, characterized in that: The hydrogel is a network structure formed by photocrosslinking a precursor solution. The precursor solution of the hydrogel includes the following components by mass fraction: 0-0.1% carboxylated multi-walled carbon nanotubes (not 0), 5%-15% methacrylate gelatin, 0.1%-1% photoinitiator, and the balance being PBS solution. The rigidity of the hydrogel can be regulated by the content of the carboxylated multi-walled carbon nanotubes.

2. The injectable hydrogel with adjustable stiffness according to claim 1, characterized in that: The carboxylated multi-walled carbon nanotubes have a diameter of 8 to 15 nm and a length of 45 to 55 μm.

3. The injectable hydrogel with adjustable stiffness according to claim 1, characterized in that: The photoinitiator is Irgacure 2925 or LAP.

4. A method for preparing an injectable hydrogel with adjustable stiffness according to any one of claims 1 to 3, characterized in that: The following steps are involved: 1) introducing the carboxylated multi-walled carbon nanotubes into a strong acid solution and performing water bath sonication to obtain a mixed solution, wherein the strong acid is a mixed acid of concentrated sulfuric acid and concentrated nitric acid; 2) diluting the mixed solution obtained in step 1) with water, then continuing to add alkaline solution to adjust the pH value of the mixed solution to 7.0, then filtering and vacuum drying the mixed solution to a constant weight to obtain surface carboxylated carbon nanotubes; 3) dispersing the carboxylated carbon nanotubes obtained in step 2) in a PBS solution containing methacrylate gelatin, adding a photoinitiator, uniformly mixing the obtained precursor solution, and then photocrosslinking to form the injectable hydrogel.

5. The method for preparing the injectable hydrogel with adjustable stiffness according to claim 4, characterized in that: The volume ratio of concentrated sulfuric acid to concentrated nitric acid in the strong acid solution in step 1) is 3:

1.

6. The method for preparing the injectable hydrogel with adjustable stiffness according to claim 4, characterized in that: The temperature of the water bath ultrasound in step 1) is 60-80° C., and the time is 4-6 hours.

7. Use of the injectable hydrogel with adjustable stiffness according to any one of claims 1 to 3 or the injectable hydrogel with adjustable stiffness prepared by the method according to any one of claims 4 to 6 in the preparation of tissue regeneration and repair materials.

8. A hydrogel scaffold for tissue repair, characterized in that: The hydrogel scaffold can be obtained by injecting a precursor solution of the injectable hydrogel with adjustable stiffness according to any one of claims 1 to 3 into a mold or a tissue defect site, and then irradiating it under ultraviolet light or visible light and performing in situ photocrosslinking.

9. The hydrogel scaffold for tissue repair according to claim 8, characterized in that: The wavelength of the ultraviolet light is 300-400 nm, the power is 15-50 W, and the irradiation time is 1-10 min.

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