A method for preparing an in situ injectable hydrogel adhesive precursor fluid that is cross-linked by visible light

Abstract

The application discloses a preparation method of visible light-induced cross-linking in-situ injectable hydrogel, and the one-step preparation process of the polymer precursor solution of the hydrogel is to blend multiple raw materials, including methacrylated gelatin, polythioctic acid, polyacrylic acid / amorphous calcium phosphate and a visible light initiation solution, and then to be injected into a target tissue in-situ under the irradiation of visible light to form a chemical / physical double-cross-linked adhesive hydrogel. The preparation process of the application is safe, simple and easy to control and large-scale industrial production, and the raw materials used are easy to obtain and low in price. The obtained hydrogel has the wet adhesion ability to the tissue surface, molecular flexibility, biodegradability, biocompatibility and high-efficiency hemostatic performance, and has the potential to be applied in non-compressive tissue wounds and other medical scenes.

Description

Technical Field

[0001] This invention relates to a medical adhesive preparation and its preparation method, specifically to a method for preparing a hydrogel precursor solution suitable for adhesion and repair of tissue damage in physiological environments. Background Technology

[0002] Hydrogels, with their adhesive properties, have become an important player in the field of tissue adhesion and have recently made significant strides in the adhesive industry. However, the water-retention capacity contributed by the unique three-dimensional network structure of hydrogels is a double-edged sword. While it provides a similar compatibility environment to surrounding tissues, it also makes hydration prone to disrupting the hydrogel's adhesive properties. Therefore, maintaining long-term stable adhesion, and ensuring the spatiotemporal consistency of degradation and repair rates remain key challenges and key considerations in the design and preparation of hydrogel adhesives.

[0003] Currently, hydrogel adhesives capable of achieving tissue adhesion mainly include pre-formed and in-situ formed types. Pre-formed hydrogels can achieve rapid adhesion to tissues within a short time using external pressure. Because it is difficult to form molecular bonds on a moist tissue surface, pre-formed hydrogels typically exhibit weak but reversible adhesive forces. To balance the interfacial adaptability of hydrogels with mechanical properties unaffected by swelling to improve adhesion durability, in-situ formed hydrogel adhesives are applied in a liquid state carrying loose polymer chains, followed by a curing process to enhance mechanical properties. The curing process, exposed to external environmental stimuli (such as light, heat, and water molecules), causes the loose polymer chains to further polymerize, increasing the polymer chain density. Therefore, in-situ formed hydrogel adhesives form molecular bonds with the tissue matrix, exhibiting strong and irreversible adhesive forces.

[0004] In-situ gelling hydrogels should possess the following basic characteristics: injectability, abundant adhesion sites, rapid curing process, and mechanical properties that match the time and space of degradation. Methacrylamide gelatin (GelMA) solutions meet the injectability requirement and, under light stimulation, gel in situ to reinforce the network structure, thus meeting the rapid curing requirement. Its controllable gelation time and tunable physical properties have attracted considerable interest. While high-energy ultraviolet light can significantly shorten gelation time, it carries the risk of oxidative DNA damage to tissues. As an alternative strategy, longer wavelength visible light can reduce tissue damage and improve cell viability. Furthermore, visible light can penetrate tissues more deeply at lower energy. The eosin Y visible light system is an FDA-approved, safe photocrosslinking system ideally suited for tissue engineering applications. During polymerization, the triplet state of eosin Y accepts hydrogen atoms from the co-initiator triethanolamine, while the comonomer N-vinylcaprolactam accelerates the gelation of methacrylamide gelatin. Although the potential of GelMA hydrogels to effectively seal tissue damage by forming stable adhesive forces during in-situ crosslinking has been demonstrated, the weak wet adhesion of pure GelMA hydrogels hinders their adhesion to moist surfaces. Furthermore, pure GelMA hydrogels are brittle due to their uneven crosslinking density and extensive entanglement. These shortcomings are expected to be mitigated by combining them with other materials.

[0005] The natural biomolecule thioctic acid (TA) contains dynamic covalent disulfide bonds and non-covalent hydrogen bonds, enabling it to construct supramolecular polymer networks through hierarchical self-assembly. The disulfide bonds of the five-membered ring in TA undergo thermally initiated ring-opening polymerization at a melting temperature of 70 °C, forming polythioctic acid (P(TA)) with a viscous fluid state. When P(TA) is cooled to physiological temperature (37 °C), the hydrogen-bonded dimerized carboxyl side chains partially crosslink with the linear covalent backbone, exhibiting a sol-gel state with high adhesiveness. Therefore, the unique properties of P(TA) promise to provide abundant adhesion sites for GelMA-based adhesive hydrogel adhesives. Furthermore, if the adhesiveness and mechanical properties of GelMA can be enhanced by adding inorganic nanoparticles, the idea of ​​producing sustainably adhesive hydrogel adhesives will be promising. Previous studies on oyster-inspired biomineralized organic-inorganic hybrid adhesives have used very small amorphous calcium carbonate (ACC) or amorphous calcium phosphate (ACP) nanoparticles as crosslinking agents to prepare and stabilize them in situ in aqueous polyacrylic acid (PAAc), thus forming a hybrid adhesive that is quite stable in an aqueous environment. However, this highly fluid and reusable adhesive lacks a curing process, and its long-term adhesion in humid environments remains a weakness. Therefore, it is hypothesized that using P(TA) and PAAc / ACP (PA) as additives to composite GelMA can improve the adhesion and mechanical toughness of visible light-induced crosslinking hydrogel adhesives. Summary of the Invention

[0006] Purpose of the invention: The purpose of this invention is to provide a method for preparing an in-situ injectable hydrogel precursor solution that is cross-linked by visible light; the second purpose of this invention is to provide a hydrogel adhesive precursor solution that has both stable tissue adhesion ability and biodegradability in a physiological environment.

[0007] Technical solution: A method for preparing a precursor solution of an in-situ injectable hydrogel adhesive with visible light-induced crosslinking, comprising the following steps:

[0008] (1) The co-initiator, comonomer, polyacrylic acid / amorphous calcium phosphate and polythioctic acid are dissolved in ultrapure water in sequence and stirred evenly to obtain a partial precursor solution ①; the co-initiator is an organic amine compound;

[0009] (2) Dissolve methacrylamide gelatin in a portion of the precursor aqueous solution ① and stir evenly under constant temperature to obtain a portion of the precursor solution ②;

[0010] (3) Dissolve the visible light initiator in part of the precursor aqueous solution ② in the dark, stir thoroughly to obtain the polymer precursor solution, and transfer it to a transferable device for storage in the dark.

[0011] In step (1), the specific preparation of polyacrylic acid / amorphous calcium phosphate is as follows: acrylic acid and thermal initiator are dissolved in ultrapure water, and reacted in a constant temperature environment for a fixed time to obtain polyacrylic acid. Then, polyacrylic acid and calcium chloride are dissolved in ultrapure water, and under stirring, disodium hydrogen phosphate solution is added dropwise at a fixed rate. After the reactants are freeze-dried and ground, polyacrylic acid / amorphous calcium phosphate powder is obtained.

[0012] The thermal initiator includes ammonium persulfate or potassium persulfate; the final concentrations of acrylic acid and the thermal initiator are 17-27% (V / V) and 0.8-1.6% (g / ml), respectively; the reaction temperature for obtaining polyacrylic acid is 60-80 °C, and the reaction time is 45-75 min; the final concentrations of polyacrylic acid and calcium chloride are 3-7% (W / W) and 1.5-2.7% (W / W), respectively; the concentration of disodium hydrogen phosphate solution is 2-3% (g / ml); the volume ratio of polyacrylic acid-calcium chloride precursor solution to disodium hydrogen phosphate solution is 6-3.75:1; and the dropping rate of disodium hydrogen phosphate solution is 1-3 mL / min.

[0013] In step (1), the specific preparation of polythioctic acid is as follows: dissolve thioctic acid in a buffer solution at a constant temperature, and then transfer the reaction solution to a constant temperature environment and let it stand for a fixed time to obtain a polythioctic acid solution.

[0014] The buffer solution is an aqueous solution of trihydroxymethylaminomethane with a concentration of 18-22% (g / ml). The reaction temperature for the dissolution process is 60-80 °C, the reaction time is 45-75 min, the reaction temperature for the standing process is 10-30 °C, and the standing time is greater than 12 h.

[0015] In step (1), the co-initiator includes any one of triethanolamine, N,N,N',N'-tetramethylethylenediamine, and tetramethylpropylenediamine, and the comonomer is N-vinylcaprolactam; the final concentrations of the co-initiator, comonomer, polyacrylic acid / amorphous calcium phosphate, and polythioctic acid are 2~3% (V / V), 1.5~2.4% (V / V), 0.8~3% (g / ml), and 10~40% (V / V), respectively.

[0016] In step (2), the specific preparation of methacrylamide gelatin is as follows: methacrylic anhydride is added dropwise to gelatin-phosphate buffer at a fixed rate, and the reaction is carried out at a constant temperature for a fixed time. After dialysis, filtration and freeze-drying, methacrylamide gelatin powder is obtained. In step (2), the final concentration of methacrylamide gelatin is 13~18% (g / ml), and the reaction temperature of part of the precursor solution ② is 37~50 ℃.

[0017] The final concentrations of gelatin and methacrylic anhydride were 8–10% (W / W) and 2.6–2.8% (W / W), respectively; the pH range of the phosphate buffer was 7.2–7.4; the dropping rate of methacrylic anhydride was 0.05–0.15 mL / min; the reaction temperature was 45–55 °C, and the reaction time was 20–24 h; the ultrapure water dialysis solution was replaced every 12 h, and dialysis was performed for 7 days at 37 °C; the pore size of the filter membrane was 0.22 µm.

[0018] In step (3), the visible light initiating solution is an aqueous solution of eosin Y with a concentration of 4~6 mmol / L, and the volume ratio of the partial precursor aqueous solution ② to the visible light initiating solution is 40~60; the transferable devices include medical devices such as syringes and minimally invasive drainage catheters.

[0019] Beneficial effects:

[0020] The polymer precursor solution prepared by this invention has temperature-responsive capabilities and can be efficiently injected in situ in a physiological environment in conjunction with various minimally invasive devices.

[0021] The hydrogel prepared by this invention has the advantages of wet adhesion to tissue surfaces, molecular flexibility, biodegradability, biocompatibility and efficient hemostatic properties.

[0022] Compared with the traditional ultraviolet light initiation system, the visible light initiation system used in this invention can reduce tissue damage and improve cell viability, making it more suitable for application in the medical field.

[0023] Compared with other similar medical adhesives, the raw materials of this invention are readily available and inexpensive, the preparation process is simple, and it is easy to operate, control, and mass-produce. It has broad application prospects in the field of tissue repair and regeneration, such as the repair of non-compression injuries. Attached Figure Description

[0024] Figure 1 This is an injection performance test of the precursor solutions of the hydrogels prepared in Examples 1-4 in a phosphate-buffered saline (PBS) environment;

[0025] Figure 2 These are photographs showing the changes in the state of the hydrogels prepared in Examples 1-4 before and after compression;

[0026] Figure 3 These are the stress-strain curves of the hydrogels prepared in Examples 2 and 4 during cyclic compression loading-unloading tests;

[0027] Figure 4 The swelling behavior of the hydrogels prepared in Examples 5 and 6 in a PBS environment;

[0028] Figure 5 The lap shear adhesion strength of the hydrogels prepared in Examples 5 and 6 to the gelatin matrix in a PBS environment;

[0029] Figure 6 The adhesion strength of the hydrogels prepared in Examples 5 and 6 to the porcine skin matrix was determined by a wound closure test in a PBS environment.

[0030] Figure 7 The effects of visible light irradiation time (b), different tissue matrices (c), and PBS soaking time (d) on the lap shear adhesion strength of the hydrogels prepared in Examples 1-4 to the gelatin matrix in PBS environment are shown in Figure 4.

[0031] Figure 8 These are photographs of the adhesion states of the hydrogels prepared in Examples 1 and 4 after they adhered to the surface of moist pigskin and were subsequently subjected to stretching, bending, twisting, PBS rinsing, and high-pressure water rinsing.

[0032] Figure 9 The images show the adhesion of the hydrogel prepared in Example 4 to the surface of various tissues in an excess PBS environment (a), and the images show the adhesion integrand after long-term immersion in PBS (b).

[0033] Figure 10The test results are for the compatibility of the hydrogels prepared in Examples 1-4 with mouse fibroblasts (L929) and human umbilical vein endothelial cells (HUVEC).

[0034] Figure 11 The degradation curves (a), weight gain curves (b), and photographs (c) of the subcutaneous implantation state of the hydrogels prepared in Examples 1 and 4 are shown.

[0035] Figure 12 The hydrogels prepared in Examples 1-4 demonstrate their in vivo hemostatic effect in an SD rat liver hemorrhage model.

[0036] Figure 13 This is a schematic diagram of the application scenarios and performance highlights of the hydrogel prepared in Example 4. Detailed Implementation

[0037] The technical solution of the present invention will be further described in detail below with reference to the accompanying drawings and embodiments.

[0038] The present invention discloses a method for preparing an in-situ injectable hydrogel adhesive with visible light-initiated crosslinking, comprising the following steps:

[0039] (1) Dissolve the co-initiator, comonomer, polyacrylic acid / amorphous calcium phosphate and polythioctic acid in ultrapure water in sequence and stir until homogeneous to obtain a partial precursor solution ①;

[0040] (2) Dissolve methacrylamide gelatin in a portion of the precursor aqueous solution ① and stir evenly under constant temperature to obtain a portion of the precursor solution ②;

[0041] (3) Dissolve the visible light initiator in part of the precursor aqueous solution ② in the dark, stir thoroughly to obtain the polymer precursor solution, and transfer it to a transferable device for storage in the dark.

[0042] (4) The polymer precursor solution is injected in situ into the damaged target site of the tissue to achieve primary adhesion. After irradiation with visible light, a stable hydrogel is obtained.

[0043] Further, in step (1), the specific preparation of polyacrylic acid / amorphous calcium phosphate is as follows: acrylic acid and a thermal initiator are dissolved in ultrapure water, and reacted at a constant temperature for a fixed time to obtain polyacrylic acid. Subsequently, polyacrylic acid and calcium chloride are dissolved in ultrapure water, and disodium hydrogen phosphate solution is added dropwise at a fixed rate while stirring. The reactants are freeze-dried and ground to obtain polyacrylic acid / amorphous calcium phosphate powder.

[0044] Furthermore, the thermal initiator includes ammonium persulfate and potassium persulfate, etc.; the final concentrations of acrylic acid and the thermal initiator are 17~27% (V / V) and 0.8~1.6% (W / V), respectively; the reaction temperature for obtaining polyacrylic acid is 60~80 °C, and the reaction time is 45~75 min; the final concentrations of polyacrylic acid and calcium chloride are 3~7% (W / W) and 1.5~2.7% (W / W), respectively; the concentration of disodium hydrogen phosphate solution is 2~3% (W / V); the volume ratio of polyacrylic acid-calcium chloride precursor solution to disodium hydrogen phosphate solution is 6~3.75; and the dropping rate of disodium hydrogen phosphate solution is 1~3 mL / min.

[0045] Furthermore, in step (1), the specific preparation of polythioctic acid is as follows: dissolve thioctic acid in a buffer solution at a constant temperature, and then transfer the reaction solution to a constant temperature environment and let it stand for a fixed time to obtain a polythioctic acid solution.

[0046] Furthermore, the buffer solution is an aqueous solution of trihydroxymethylaminomethane with a concentration of 18-22% (W / V), the reaction temperature of the dissolution process is 60-80 °C, the dissolution reaction time is 45-75 min, the reaction temperature of the standing process is 10-30 °C, and the standing reaction time is greater than 12 h.

[0047] Furthermore, in step (1), the co-initiator includes triethanolamine, N,N,N',N'-tetramethylethylenediamine, tetramethylpropylenediamine, etc., and the comonomer is N-vinylcaprolactam; the final concentrations of the co-initiator, comonomer, polyacrylic acid / amorphous calcium phosphate and polythioctic acid are 2~3% (V / V), 1.5~2.4% (V / V), 0.8~3% (W / V) and 10~40% (V / V), respectively.

[0048] Furthermore, in step (2), the specific preparation of the methacrylated gelatin is as follows: methacrylic anhydride is added dropwise to the gelatin-phosphate buffer at a fixed rate, and the reaction is carried out at a constant temperature for a fixed time. After dialysis, filtration, and freeze-drying, the reactants are used to obtain methacrylated gelatin powder. In step (2), the final concentration of the methacrylated gelatin is 13-18% (W / V), and the reaction temperature of part of the precursor solution ② is 37-50 °C.

[0049] Furthermore, the final concentrations of the gelatin and methacrylic anhydride are 8-10% (W / W) and 2.6-2.8% (W / W), respectively; the pH range of the phosphate buffer is 7.2-7.4; the dropping rate of the methacrylic anhydride is 0.05-0.15 mL / min; the reaction temperature is 45-55 °C, and the reaction time is 20-24 h; the ultrapure water dialysis solution is replaced every 12 h, and dialysis is performed for 7 days at 37 °C; the pore size of the filter membrane is 0.22 µm.

[0050] Furthermore, in step (3), the visible light initiating solution is an aqueous solution of eosin Y with a concentration of 4~6 mmol / L, and the volume ratio of the partial precursor aqueous solution ② to the visible light initiating solution is 40~60; the transferable devices include medical devices such as syringes and minimally invasive drainage catheters.

[0051] Furthermore, in step (4), the brightness of the visible light device is 1000~6000 lm, the distance between the device and the adhered tissue is 1~5 cm, and the irradiation time is 1~10 min.

[0052] Preparation principle of this invention: The polymer precursor solution in this invention is prepared through a simple procedure, consisting of a physically cross-linked polymer network embedded with inorganic nanoparticles. This physically cross-linked polymer network is constructed by mixing GelMA, P(TA), PA, and a visible light initiator, and is injectable at physiological temperatures. Hydrogen bonds and electrostatic interactions exist between the three main polymer components, mitigating the risk of the precursor solution slipping from the target site before gelation. The precursor solution is injected in situ into the damaged tissue target site via a delivery system, such as a syringe or minimally invasive catheter. Thanks to the synergistic effect of multiple non-covalent interactions, including hydrogen bonds and electrostatic interactions, the precursor solution can temporarily adhere to the target site. Subsequently, upon irradiation with visible light, the double bonds in GelMA break, initiating inter-chain cross-linking, and the precursor solution rapidly solidifies to form a chemically / physically cross-linked hydrogel. This improves mechanical properties, allows for long-term, firm adhesion to the tissue surface, and provides biodegradable properties. This safe combination of injectable strategy and polymer system holds promise for applications in tissue repair and regeneration, such as the repair of non-compression injuries.

[0053] Example 1

[0054] (1) In a 50 °C water bath environment, 167 μL of triethanolamine, 121 μL of N-vinylcaprolactam and 1 g of methacrylamide gelatin were dissolved in 5 mL of ultrapure water and stirred until homogeneous.

[0055] (2) Add 100 μL of 5 mM eosin Y solution to the above solution in the dark, stir thoroughly to obtain polymer precursor solution, transfer to syringe and store in the dark;

[0056] (3) The polymer precursor solution was incubated at 37 °C for 10 min and then injected in situ into the target site. After irradiating the target site with an LED tube with a brightness of 1000 lm at a distance of 2 cm for 10 min, GelMA hydrogel was obtained.

[0057] Example 2

[0058] (1) Dissolve 167 μL of triethanolamine, 121 μL of N-vinylcaprolactam and 1.25 mL of polythioctic acid in 3.75 mL of ultrapure water and stir until homogeneous to obtain a partial precursor solution ①.

[0059] (2) Dissolve 1 g of methacrylamide gelatin in a water bath at 50 °C in a portion of the precursor aqueous solution ① to obtain a portion of the precursor solution ②;

[0060] (3) Add 100 μL of 5 mM eosin Y solution to part of the precursor aqueous solution ② in the dark, stir thoroughly to obtain polymer precursor solution, transfer to syringe and store in the dark;

[0061] (4) The polymer precursor solution was incubated at 37 °C for 10 min and then injected in situ into the target site. After irradiating the target site with an LED tube with a brightness of 1000 lm at a distance of 2 cm for 10 min, GelMA / P(TA) hydrogel was obtained.

[0062] Example 3

[0063] (1) Dissolve 167 μL of triethanolamine, 121 μL of N-vinylcaprolactam and 100 mg of polyacrylic acid / amorphous calcium phosphate in 5 mL of ultrapure water and stir until homogeneous to obtain a partial precursor solution ①.

[0064] (2) Dissolve 1 g of methacrylamide gelatin in a water bath at 50 °C in a portion of the precursor aqueous solution ① to obtain a portion of the precursor solution ②;

[0065] (3) Add 100 μL of 5 mM eosin Y solution to part of the precursor aqueous solution ② in the dark, stir thoroughly to obtain polymer precursor solution, transfer to syringe and store in the dark;

[0066] (4) The polymer precursor solution was incubated at 37 °C for 10 min and then injected in situ into the target site. After irradiating the target site with an LED tube with a brightness of 1000 lm at a distance of 2 cm for 10 min, the GelMA / PA hydrogel was obtained.

[0067] Example 4

[0068] (1) Dissolve 167 μL of triethanolamine, 121 μL of N-vinylcaprolactam, 100 mg of polyacrylic acid / amorphous calcium phosphate and 1.25 mL of polythioctic acid in 3.75 mL of ultrapure water and stir until homogeneous to obtain a partial precursor solution ①.

[0069] (2) Dissolve 1 g of methacrylamide gelatin in a water bath at 50 °C in a portion of the precursor aqueous solution ① to obtain a portion of the precursor solution ②;

[0070] (3) Add 100 μL of 5 mM eosin Y solution to part of the precursor aqueous solution ② in the dark, stir thoroughly to obtain polymer precursor solution, transfer to syringe and store in the dark;

[0071] (4) The polymer precursor solution was incubated at 37 °C for 10 min and then injected in situ into the target site. After irradiating the target site with an LED tube with a brightness of 1000 lm at a distance of 2 cm for 10 min, GelMA / P(TA) / PA (GelMA / P(TA)) was obtained. 25 / PA2) hydrogel.

[0072] Example 5

[0073] Five parallel experiments were designed, with the specific steps the same as in Example 4. The difference was that the volumes of polythioctic acid added in step (1) were 0 mL, 0.625 mL, 1.25 mL, 1.875 mL, and 2.5 mL, respectively, and the corresponding ultrapure water contents were 5 mL, 4.375 mL, 1.25 mL, 3.125 mL, and 2.5 mL, respectively. The hydrogels were represented in the figure as GelMA / P(TA)0 / PA2, GelMA / P(TA), etc. 12.5 / PA2, GelMA / P(TA) 25 / PA2, GelMA / P(TA) 37.5 / PA2 and GelMA / P(TA) 50 / PA2.

[0074] Example 6

[0075] Five parallel experiments were designed, with the specific steps the same as in Example 4. The difference was that the mass of polyacrylic acid / amorphous calcium phosphate added in step (1) was 0 mg, 50 mg, 100 mg, 150 mg and 200 mg, respectively, and each hydrogel was represented as GelMA / P(TA) in the figure. 25 / PA0、GelMA / P(TA) 25 / PA1, GelMA / P(TA) 25 / PA2, GelMA / P(TA) 25 / PA3 and GelMA / P(TA) 25 / PA4.

[0076] Example 7 Preparation of polyacrylic acid / amorphous calcium phosphate

[0077] (1) Dissolve 3.46 mL of acrylic acid and 0.19 g of ammonium persulfate in 11.54 mL of ultrapure water and react in a water bath at 70 °C for 1 h to obtain polyacrylic acid hydrogel;

[0078] (2) The above polyacrylic acid hydrogel was dissolved in 300 mL of calcium chloride aqueous solution (0.2 M) by vigorous stirring, and then 60 mL of disodium hydrogen phosphate solution (0.17 M) was added dropwise to the above solution at a rate of 2 mL / min.

[0079] (3) During the reaction, the solution becomes turbid and a white viscous substance gradually accumulates around the stirring rod. The turbid reaction solution is continuously stirred for 6 hours until the solution becomes clear. After freeze-drying and grinding the viscous substance, polyacrylic acid / amorphous calcium phosphate powder can be obtained.

[0080] The injectability of the precursor solutions in PBS in each embodiment was initially evaluated using a 27 G syringe needle. Figure 1 As can be observed from the images, the precursor solutions of the hydrogels prepared in Examples 1-3 were difficult to mold after being extruded into PBS using a needle. Furthermore, the "SEU" lettering written by injection disintegrated after being soaked in PBS for 1 hour, becoming blurred and illegible. Conversely, thanks to the non-covalent interactions between the three main components and the higher polymer concentration, the "SEU" lettering extruded using the precursor solution of the hydrogel prepared in Example 4 remained intact. This demonstrates that the hydrogel precursor solution of Example 4, while maintaining injectability, is capable of avoiding the risk of being washed away by physiological fluids and thus failing to exert its adhesive effect. Figure 2 The P(TA)-containing composite gels (Examples 2 and 4) were shown to rapidly recover their original shape after being compressed to ~90% strain. Furthermore, after undergoing five compression loading-unloading cycles, they exhibited only a slight loss of toughness, manifested as a slight change in the stress-strain curve. Figure 3 ).

[0081] The swelling curves of the hydrogels prepared in Examples 5 (different P(TA) concentrations) and 6 (different PA concentrations) in the PBS environment show that the hydrogel system basically reaches swelling equilibrium after 1 day. Figure 4The hydrogel of Example 1, relying on a stable covalent crosslinking network, exhibited a near-zero swelling ratio. The hydrogel of Example 2 exhibited the highest swelling ratio due to the combined effects of a lower degree of covalent crosslinking and abundant carboxyl groups. Amorphous calcium phosphate complexed most of the carboxyl groups on the polyacrylic acid chains, resulting in a swelling ratio of only ~50% for the hydrogel of Example 3. The physical interactions between P(TA) and PA, including hydrogen bonding and electrostatic interactions, induced a lower swelling ratio in the hydrogel of Example 4 compared to that of Example 2.

[0082] The adhesion strength of the hydrogels prepared in Examples 5 and 6 on gelatin or pigskin substrates in PBS was determined according to the overlap shear adhesion test (ASTM F2255) and the wound closure test (ASTM F2458-05). Figure 5 6). After the same amount of visible light initiation, the adhesion strength of the composite hydrogel increased with increasing P(TA) concentration, which is related to the higher precursor viscosity and more abundant adhesion groups. The hydrogels of the examples with excessively high PA concentrations exhibited cohesive forces higher than adhesion forces, resulting in lower adhesion strength to the substrate. The adhesion strength of the hydrogels of the examples reached its maximum critical value when the PA concentration was 2 wt%.

[0083] A series of more in-depth adhesion behavior evaluations were conducted on the hydrogel of Example 4. Due to a richer array of surface adhesion functional groups and a more harmonious ratio of covalent / non-covalent crosslinked networks, the hydrogel of Example 4 exhibited higher adhesion strength than the hydrogels of Examples 1-3. Figure 7 a). Example 4: The lap shear adhesion strength of the hydrogel exceeds 100 kPa after 2 minutes of visible light irradiation and stabilizes after 10 minutes of irradiation. The shorter adhesion time facilitates adhesion application in actual surgical procedures. Figure 7 b). Example 4: The hydrogel also exhibits non-specific adhesion to different biological tissue matrices ( Figure 7 c). After being stretched, bent, deformed, and even rinsed with PBS and high-pressure tap water, the hydrogel of Example 4 remained firmly adhered to the moist pigskin surface ( Figure 8 However, after the same operation, it was found that the hydrogel of Example 1 ruptured and peeled off from the pigskin surface. Figure 8 Similar to most hydrogel adhesives, the lap joint between the hydrogel and gelatin matrix in Example 4 tends to decrease in adhesion strength with prolonged immersion time in PBS. Figure 7d). Furthermore, performing adhesion operations on hydrogel adhesives in excessive liquid environments is often extremely challenging. Excitingly, in Example 4, the hydrogel precursor solution, after in situ injection into the surface of biological tissue in excess PBS, followed by 10 minutes of visible light irradiation to initiate gelation, achieved robust adhesion of the hydrogel to various tissue surfaces. Figure 9 a). After prolonged immersion in PBS, the sustained adhesion time of the hydrogel in Example 4 to tissue surfaces varied depending on the type of tissue tested. Its deadhesion times on the skin, myocardium, stomach, liver, and large intestine were 11, 17, 20, 21, and >30 days, respectively. Figure 9 b). Example 4: The strong adhesion of the hydrogel to the surface of biological tissues may be attributed to the synergistic effect of multiple non-covalent interactions, including hydrogen bonding and electrostatic interactions.

[0084] The cell compatibility of the extracts from the hydrogels of Examples 1-4 with mouse fibroblasts (L929) and human umbilical vein endothelial cells (HUVECs) was evaluated using the CCK-8 assay. The results showed that the extracts from all four hydrogels were beneficial to the growth of both cell types. Furthermore, the cell proliferation rate in the Example 4 hydrogel group was superior to that in the Example 1 hydrogel group, a phenomenon particularly pronounced for HUVEC cells. Figure 10 The in vivo biocompatibility of the hydrogels from Examples 2 and 4 was further examined through subcutaneous implantation experiments over 30 days. The hydrogel from Example 1 showed minimal degradation (approximately ~10%) after 30 days of subcutaneous implantation, while the hydrogel from Example 4 completely degraded within 30 days. Commercially available fibrin gels, being composed solely of a physically cross-linked network, exhibited the fastest degradation rate; completely degraded samples were detectable after 15 days. Figure 11 a). The body weight of all experimental groups showed a steady, time-dependent increase with no significant differences among them. Figure 11 (b) and no obvious edema, congestion or other inflammatory reactions were observed at the implantation site macroscopically. Figure 11 c). Therefore, it can be concluded that the hydrogel prepared by this invention, while possessing in vivo degradability, induces a comparablely minimal inflammatory response in vivo. Overall, comprehensive in vivo and in vitro evaluations of the hydrogel indicate that the hydrogel prepared by this invention exhibits good biodegradability and biocompatibility.

[0085] Finally, the in vivo hemostatic effect of the hydrogels of Examples 1-4 was evaluated by constructing a liver injury and bleeding model in SD rats. After in vivo induction of rat liver injury, the hydrogel precursor fluid was immediately transferred to the bleeding site via syringe, followed by continuous irradiation with visible light for 2 minutes to induce a gelation reaction, thereby achieving hemostasis. Since the commercially available fibrin gel is not a photo-initiated cross-linked hydrogel system, visible light irradiation was not performed after applying this gel. The results showed that compared with the untreated control group, both the commercially available hemostatic gel and the hydrogels of each example showed significantly improved hemostatic effect, manifested as lower blood loss. No significant differences were detected among the hydrogels of each example, and the bleeding volume of the three hydrogels other than Example 1 was significantly lower than that of the commercially available hemostatic gel. The lowest bleeding volume was obtained from the filter paper of the hydrogel of Example 4, only 77.73 ± 18.05 mg, indicating the best hemostatic effect. The rapid hemostatic ability of the hydrogel prepared in this invention can be attributed to its rapid and stable wound closure in moist environments. When a non-flowing hydrogel precursor fluid, mediated by physical interactions and cross-linking, is applied to a bleeding site, its surface adhesive functional groups rapidly form non-covalent intermolecular bonds with the tissue surface, such as hydrogen bonds and electrostatic forces. This resulting adhesive layer acts as a primary physical barrier, preventing blood leakage for a short period. Subsequently, with the synergistic effect of the initiation system, visible light irradiation induces rapid cross-linking of methacryloyl gelatin, thereby constructing a chemically / physically cross-linked hydrogel with molecular flexibility, which enhances the stability of the hemostatic barrier. Therefore, the safe and efficient in vivo hemostatic effect exhibited by the hydrogel prepared in this invention holds promise for playing a role in the emergency treatment of non-compression injuries.

Claims

1. A method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution, characterized in that, Includes the following steps: (1) The co-initiator, comonomer, polyacrylic acid / amorphous calcium phosphate and polythioctic acid are dissolved in ultrapure water in sequence and stirred evenly to obtain a first precursor aqueous solution; the co-initiator is an organic amine compound; the comonomer is N-vinylcaprolactam; the amorphous calcium phosphate is nanoparticles; the final concentrations of the co-initiator, comonomer, polyacrylic acid / amorphous calcium phosphate and polythioctic acid are 2-3% by volume, 1.5-2.4% by volume, 0.8-3% by mass / volume in g / ml and 10-40% by volume, respectively; (2) Dissolve methacrylated gelatin in the first precursor aqueous solution and stir evenly at 37~50 ℃ to obtain a pre-reaction solution; the specific preparation of the methacrylated gelatin is as follows: add methacrylic anhydride dropwise to gelatin-phosphate buffer at a rate of 0.05~0.15 mL / min, and react at 45~55 ℃ for 20~24 h. After dialysis, filtration and freeze-drying, the reactants are used to obtain methacrylated gelatin powder; the final concentration of methacrylated gelatin is 13~18% by mass / volume in g / ml. (3) Dissolve the visible light initiator in the pre-reaction solution in the dark, stir thoroughly to obtain the polymer precursor solution, and transfer it to a transferable device for storage in the dark.

2. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 1, characterized in that: The specific preparation of polyacrylic acid / amorphous calcium phosphate in step (1) is as follows: acrylic acid and thermal initiator are dissolved in ultrapure water and reacted in a constant temperature environment for a fixed time to obtain polyacrylic acid. Then, polyacrylic acid and calcium chloride are dissolved in ultrapure water and disodium hydrogen phosphate solution is added dropwise at a fixed rate under stirring. After freeze-drying and grinding, polyacrylic acid / amorphous calcium phosphate powder is obtained. The reaction temperature for obtaining polyacrylic acid is 60~80℃ and the reaction time is 45~75 min. The dropping rate of disodium hydrogen phosphate solution is 1~3 mL / min.

3. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 2, characterized in that: The thermal initiator includes ammonium persulfate or potassium persulfate; the final concentrations of acrylic acid and the thermal initiator are 17-27% by volume and 0.8-1.6% by mass / volume (g / ml), respectively; the final concentrations of polyacrylic acid and calcium chloride are 3-7% by mass percentage and 1.5-2.7% by mass percentage, respectively; the concentration of disodium hydrogen phosphate solution is 2-3% by mass / volume (g / ml); and the volume ratio of polyacrylic acid-calcium chloride precursor solution to disodium hydrogen phosphate solution is 6-3.75:

1.

4. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 1, characterized in that: In step (1), the specific preparation of polythioctic acid is as follows: dissolve thioctic acid in buffer solution at 60~80 ℃ for 45~75 min, then transfer the reaction solution to an environment of 10~30 ℃ and let it stand to obtain polythioctic acid solution. The standing reaction time is greater than 12 h.

5. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 4, characterized in that: The buffer solution is an aqueous solution of trihydroxymethylaminomethane with a mass-volume concentration of 18-22% (g / ml).

6. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 1, characterized in that: In step (1), the co-initiator includes any one of triethanolamine, N,N,N',N'-tetramethylethylenediamine, and tetramethylpropylenediamine.

7. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 6, characterized in that: The final concentrations of the gelatin and methacrylic anhydride were 8-10% and 2.6-2.8% by mass, respectively; the pH range of the phosphate buffer was 7.2-7.4; the ultrapure water dialysis solution was replaced every 12 h, and dialysis was performed for 7 days at 37 ℃; the pore size of the filter membrane was 0.22 µm.

8. The method for preparing a visible light-initiated crosslinking in-situ injectable hydrogel adhesive precursor solution according to claim 1, characterized in that: In step (3), the visible light initiating solution is an aqueous solution of eosin Y with a concentration of 4~6 mmol / L, and the volume ratio of the pre-reaction solution to the visible light initiating solution is 40~60:1; the transferable devices include syringes, minimally invasive drainage catheters and other medical devices.

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