Medical waterproof adhesive as well as preparation method and application thereof

Abstract

The invention provides a medical waterproof adhesive as well as a preparation method and application thereof. The preparation method of the medical waterproof adhesive comprises the following steps: firstly preparing solid gel, and then adjusting the rheological property of the solid gel to obtain the viscoelastic fluid waterproof adhesive. The preparation method of the adhesive provided by the invention is simple and convenient to operate, controllable in process, high in repeatability and low in cost. The prepared waterproof adhesive is stable in performance, shows good injectability and spreadability at normal temperature, and is convenient for surgical operation; in a wet ocular surface environment, a pre-formed network can effectively resist immediate flushing and dilution of body fluids such as tears, aqueous humor and blood, so that stable residence and accurate positioning on the surface of a target tissue are ensured; under the irradiation of a light source with specific wavelength, the gel is subjected to polymerization reaction, a compact and firm secondary cross-linked network is quickly formed under a wet condition, and stable and reliable in-situ solidification and tissue adhesion are realized.

Description

Technical Field

[0001] This invention belongs to the field of biomedical materials, specifically relating to a medical waterproof adhesive, its preparation method, and its application. Background Technology

[0002] Over the past few decades, microsurgery has made continuous progress in terms of operational precision and instrument design, but tissue apposition still heavily relies on traditional suturing methods. Especially in complex surgeries such as corneal transplantation, dozens of microsurgical sutures are often required to ensure stable graft fixation and wound closure. Microsurgical suturing not only demands high levels of expertise from both the surgical equipment and the operator, but also easily leads to postoperative complications, including suture loosening or breakage, infectious keratitis, aseptic infiltration, corneal neovascularization, and iatrogenic high astigmatism. Therefore, the development of novel bio-adhesives or sutureless repair materials that can achieve reliable tissue apposition and reduce reliance on sutures has become an urgent clinical need.

[0003] To overcome these limitations, various medical adhesives have been developed and applied to soft tissue repair. Currently, commonly used tissue adhesives in clinical practice mainly include cyanoacrylate glue, fibrin glue, and polyethylene glycol gels. Among them, cyanoacrylate glue has the advantages of simple operation and rapid curing, but the chemical monomers released when polymerization is incomplete may irritate tissues, leading to scar hyperplasia and neovascularization, making it particularly unsuitable for ophthalmic applications. Although fibrin glue and polyethylene glycol gel exhibit better biocompatibility, they still have problems such as short curing time windows, insufficient adhesion in aqueous environments, the need for on-site mixing of multiple components, and low precision, which limit their clinical applicability and ease of operation.

[0004] To meet the urgent clinical need for highly efficient tissue adhesive materials in ophthalmic surgery, various novel in-situ curable adhesive hydrogel materials have been extensively researched and developed. For example, patent document CN115554461A discloses a strategy of grafting methacrylamide onto gelatin molecular chains through photo-initiated free radical polymerization to prepare a highly transparent ocular surface adhesive that can be cured by blue light; patent document CN116099034A reports a photocurable bioadhesive based on a composite system of methacrylamide gelatin and aldehyde-modified F127, combined with riboflavin photosensitizer, to enhance material toughness and tissue adhesion strength; patent document CN114159625A further proposes a three-dimensional composite hydrogel formed by interpenetrating alginate hydrogel and decellularized matrix hydrogel crosslinked with transglutaminase, for corneal tissue repair.

[0005] It is worth noting that the adhesives mentioned in the aforementioned patent documents are all liquid solutions. However, in many moist surgical environments containing bodily fluids, irrigation solutions, or exudates, such as ophthalmic surgery, orthopedic arthroscopic surgery, urological endoscopic surgery, and minimally invasive abdominal or thoracic surgery, the surgical area is often in a moist, dynamic liquid environment. Taking typical ophthalmic surgery as an example, during the operation, physiological saline is continuously used to maintain the moisture of the ocular surface to prevent tissue exposure damage; the intraocular space is entirely composed of aqueous humor and vitreous humor; in addition, ocular surface surgery is often accompanied by bleeding. Under such moist, dynamic liquid conditions, existing liquid adhesives are easily diluted by irrigation solutions or bodily fluids, leading to a decrease in the concentration of effective ingredients, a reduction in cross-linking efficiency, or even complete loss of adhesive function. At the same time, due to the high fluidity of liquids, adhesives are prone to runoff or spread from the target area before curing, causing complications such as adhesion to non-target tissues and blockage of important physiological cavities (such as lacrimal ducts, sinus ostia, and small blood vessel ends). In ophthalmic surgery, if liquid adhesives flow into the anterior chamber or vitreous cavity, they may trigger an inflammatory response; if they solidify and remain near the angle of the anterior chamber, they may interfere with the aqueous humor circulation pathway, potentially inducing secondary glaucoma. In other intracavitary surgeries, accidental diffusion or loss of adhesives can also lead to tissue adhesions, foreign body reactions, or functional impairment.

[0006] Therefore, the insufficient anti-dilution and positioning retention of liquid adhesives in moist surgical environments can easily lead to adhesive failure or complications. This has become a key technical bottleneck restricting their clinical translation and widespread application in different liquid surgical scenarios. Summary of the Invention

[0007] To overcome the shortcomings of the prior art, in a first aspect, the present invention provides a method for preparing a medical waterproof adhesive, which first prepares a solid gel through chemical crosslinking, and then adjusts the rheological properties of the solid gel through hydrolysis, including the following steps:

[0008] Solid gels with a three-dimensional cross-linked network with amide bonds as connection points were obtained by using coupling agents to catalyze collagen materials;

[0009] The solid gel is placed in an aqueous environment to partially break the amide bonds inside the solid gel, thereby adjusting the viscosity and obtaining a viscoelastic fluid.

[0010] Under light irradiation, the viscoelastic fluid forms a hydrophilic polymer network on a chemical cross-linking network, resulting in a solid waterproof adhesive.

[0011] Preferably, the preparation method of the medical waterproof adhesive includes the following steps:

[0012] Step 1) Prepare a composite crosslinking solution, the components of which include: 0.1-0.4% (w / v) photoinitiator, no more than and including 20% ​​(w / v) collagen material, 5-30% (v / v) methacrylamide derivative and 0.05-0.5% (w / v) coupling agent; more preferably, the components include: 0.2-0.3% (w / v) photoinitiator, 5-20% (w / v) collagen material, 5-20% (v / v) methacrylamide derivative and 0.05-0.5% (w / v) coupling agent;

[0013] Step 2) The composite cross-linking solution is placed at the reaction temperature to react and obtain a solid gel with a three-dimensional chemical cross-linking network with amide bonds as the connection points;

[0014] Step 3) The solid gel is placed in a high-temperature aqueous environment for hydrolysis. When the viscosity drops to 2000-20000 mPa·s, a viscoelastic fluid-like waterproof adhesive is obtained; the high temperature refers to 60-150°C.

[0015] Preferably, the collagen material is gelatin or recombinant human collagen.

[0016] Preferably, the coupling agent is a compound or composition capable of activating the carboxyl group and causing it to form an amide bond with the amine group.

[0017] Preferably, the coupling agent is a combination of carbodiimide and N-hydroxy compound; or the coupling agent is a triazine salt coupling agent; or the coupling agent is a ureonium / phosphonium salt coupling agent.

[0018] More preferably, the carbodiimide is selected from one or more of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide, N,N'-dicyclohexylcarbodiimide, N,N'-diisopropylcarbodiimide, and 1-cyclohexyl-2-morpholinoethylcarbodiimide; the N-hydroxy compound is selected from one or more of N-hydroxysuccinimide and sulfonyl-N-hydroxysuccinimide.

[0019] More preferably, the triazineonium salt coupling agent is 4-(4,6-dimethoxy-1,3,5-triazine-2-yl)-4-methylmorpholinonium salt.

[0020] More preferably, the ureonium / phosphonium salt coupling agent is selected from one or more of O-(7-azabenzotriazol-1-yl)-N,N,N',N'-tetramethylurea hexafluorophosphate, benzotriazol-1-yl-oxytripyrrolylphosphine hexafluorophosphate, or (benzotriazol-1-yloxy)tri(dimethylamino)phosphonium hexafluorophosphate.

[0021] Preferably, the photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP), sodium phenyl-2,4,6-trimethylbenzoylphosphonate (NAP), or a combination thereof.

[0022] Preferably, the methacrylamide derivative is selected from one or more of polyethylene glycol diacrylate (PEGDA), methacrylamide gelatin (GelMA), methacrylamide chitosan (CsMA), methacrylamide chondroitin sulfate (ChsMA), and methacrylamide dextran (DexMA).

[0023] Preferably, the reaction temperature refers to standing at 10-34°C; the high temperature refers to 85-120°C.

[0024] In a second aspect, the present invention provides a waterproof adhesive prepared by the above-described method. The waterproof adhesive possesses waterproof properties, dilution resistance, and erosion resistance; its comprehensive performance parameters include, but are not limited to, light transmittance of ≥80%; refractive index of 1.33-1.40; volume swelling ratio of 1.0-5.0; underwater oscillation mass loss rate of less than 10% after 1 hour; migration rate of less than 10% after 1 minute of 45° tilting; shear strength of 45-55 kPa; tensile strength of 25-40 kPa; and interfacial toughness of 100-130 Jm. -2 .

[0025] In a third aspect, the present invention provides the application of the above-mentioned waterproof adhesive in the preparation of medical devices for use in humid environments. These humid environments include, but are not limited to, medical scenarios such as wound care, mucosal protection, laparoscopy, endoscopy, and surgery.

[0026] Preferably, the waterproof adhesive is used in ophthalmic surgery.

[0027] Beneficial effects:

[0028] The adhesive preparation method provided by this invention is simple to operate, has controllable process, high repeatability, and low cost.

[0029] The waterproof adhesive provided by this invention has stable performance and exhibits good injectability and spreadability at room temperature, which facilitates surgical operation. In the moist ocular surface environment, its pre-formed covalent cross-linked network can effectively resist the immediate flushing and dilution of bodily fluids such as tears, aqueous humor and blood, ensuring its stable residence and precise positioning on the target tissue surface.

[0030] The waterproof adhesive provided by this invention can undergo a highly efficient photo-initiated free radical polymerization reaction of the methacrylate groups in the gel system under the irradiation of a specific wavelength light source (such as 405nm or 365nm blue light), and rapidly form a dense and firm secondary cross-linked network under humid conditions, thereby achieving stable and reliable in-situ curing and tissue adhesion.

[0031] This invention effectively solves the problems of traditional liquid adhesives being easily diluted, difficult to position, and having poor curing effect in moist surgical environments. It provides a safe, efficient, and easy-to-operate sutureless adhesive and wound closure solution for surgeries performed on moist tissue surfaces or in the presence of body fluids. Attached Figure Description

[0032] Figure 1 The anti-dilution ability of the adhesive prepared in Example 1 and the control group sample in PBS solution at 37°C is shown, wherein: A is control group 1; B is control group 2; and C is Example 1.

[0033] Figure 2 The adhesive prepared in Example 1 and the control sample are shown to maintain their position on a 45° inclined plane, where: A is Example 1; B is control 3.

[0034] Figure 3 The adhesive strength of the adhesive prepared in Example 1 and the control group adhesive are shown, where: A is the peel test; B is the shear test; and C is the tensile test.

[0035] Figure 4 The images show live and dead fluorescence staining of corneal epithelial and stromal cell lines and the statistical results of cell viability, where: A represents corneal epithelial cells; B represents corneal stromal cells; and C represents a statistical graph of cell viability.

[0036] Figure 5 The underwater tissue adhesion effect of the waterproof adhesive is shown. In A, the operation process of applying the waterproof gel to a small intestinal mucosal tissue slice and then curing it in an underwater environment is shown. In B, after light curing, two small intestinal mucosal tissue slices can be lifted simultaneously after being bonded by the gel, demonstrating its underwater tissue adhesion performance.

[0037] Figure 6 The images show follow-up photos taken on the day of surgery and 7 days after surgery for repairing corneal stromal defects using the adhesive prepared in Example 1 and the control adhesive.

[0038] Figure 7 Gross images of the sutureless and fused conjunctival grafts are shown at 7 and 14 days post-surgery. White arrows indicate the graft boundaries, and black arrows indicate perigraft bleeding following suture removal one week post-surgery.

[0039] Figure 8 Slit-lamp micrographs of two patients 1 day and 7 days after sutureless conjunctival transplantation are shown. Detailed Implementation

[0040] This invention provides a waterproof gel (adhesive) suitable for humid / underwater environments. The waterproof gel is prepared through a stepwise construction and precise control process, namely, first constructing an insoluble covalent skeleton, and then adjusting the macroscopic rheological properties through controllable hydrolysis. Specifically, it includes the following two steps:

[0041] Construction of an insoluble gel framework: Coupling agents are used to catalyze the formation of solid gels with stable covalent amide bond networks from collagen-based materials (a class of polymers based on natural proteins or their derivatives). This step constructs a continuous, complete, and water-insoluble three-dimensional cross-linked network at the molecular level. This network is the structural basis for the material's permanent "waterproof," "solution-resistant," and "dilution-resistant" properties, which are maintained by adhesives in subsequent processing and applications.

[0042] High-temperature hydrothermal regulation of rheological properties: The above-mentioned solid gel is placed in a high-temperature aqueous environment (e.g., 60-120°C) for viscosity adjustment to obtain a viscoelastic fluid. This step induces controlled, partial hydrolytic breakage of amide bonds under high-temperature hydrothermal conditions, thereby systematically reducing the crosslinking density without disrupting the overall network continuity, thus endowing the adhesive with deformability. By precisely controlling temperature and time, the macroscopic state of the material can be repeatedly and accurately controlled from a solid elastomer to a fluid with appropriate viscoelasticity, facilitating subsequent minimally invasive delivery and application.

[0043] The two steps described above, while ensuring that the material does not dissolve, simultaneously meet the clinical requirements for injectability and deformability, achieving a balance between the adhesive's resistance to dilution and its operability in application.

[0044] Under irradiation with a specific wavelength light source (such as 405nm), the viscoelastic fluid obtained after rheological regulation rapidly forms a second dense and hydrophilic polymer network. This network interpenetrates with the aforementioned partially hydrolyzed collagen covalent network, together forming a stable interpenetrating polymer network, thereby achieving rapid and firm in-situ curing and adhesion on the surface of moist tissue.

[0045] Specifically, the waterproof adhesive provided by this invention has suitable adhesion and strong erosion resistance, and the preparation steps include:

[0046] Step 1) Prepare a composite crosslinking solution, the main components of which include: photoinitiator, no more than 20% (w / v) collagen material, 5-30% (v / v) methacrylamide derivative and 0.05-0.5% (w / v) coupling agent;

[0047] Step 2) Constructing an insoluble gel framework: The composite cross-linking solution is placed at a suitable temperature (such as below 60°C) for reaction, and a solid gel with a three-dimensional cross-linked network with amide bonds as connection points is formed by the catalysis of a coupling agent.

[0048] Step 3) High-Temperature Hydrothermal Regulation of Rheological Properties: The solid gel obtained in Step 2 is subjected to high-temperature hydrothermal treatment. Its rheological properties are monitored in real time until its dynamic viscosity decreases and stabilizes within the range of 2000-20000 mPa·s, thus obtaining a viscoelastic fluid. This step precisely "softens" the insoluble framework through controlled hydrolysis, transforming it into a viscoelastic fluid that can be smoothly dispensed via a syringe. At this viscosity range, the adhesive can be smoothly delivered through minimally invasive instruments (such as 22G-27G needles) and resists dilution by tissue fluid before solidification.

[0049] Those skilled in the art can adjust the viscosity to other ranges as needed. The preferred viscosity of this invention is 2000-20,000 mPa·s to maintain sparse covalent crosslinking points and highly entangled polymer chains within the adhesive. This structure imparts significant viscosity and intrinsic strength, effectively resisting dilution and shear forces from water flow when exposed to a liquid environment, preventing it from being washed away, while also exhibiting good injectability.

[0050] Preferably, the photoinitiator is selected from one or more of lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP) and sodium phenyl-2,4,6-trimethylbenzoylphosphonate (NAP).

[0051] Preferably, the collagen material includes, but is not limited to, gelatin, recombinant human collagen, etc.

[0052] Preferably, the methacrylamide derivative is selected from one or more of polyethylene glycol diacrylate (PEGDA), methacrylamide gelatin (GelMA), methacrylamide chitosan (CsMA), methacrylamide chondroitin sulfate (ChsMA), and methacrylamide dextran (DexMA).

[0053] Preferably, the coupling agent refers to an activation system / combination used to covalently link carboxyl and amine groups to form an amide bond. Its core function is to convert the carboxyl group into a highly reactive intermediate. This activation system includes, but is not limited to, any of the following compounds or combinations thereof: (a) a combination system of carbodiimide and N-hydroxy compound, wherein the carbodiimide is selected from 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), N,N'-dicyclohexylcarbodiimide (DCC), N,N'-diisopropylcarbodiimide (DIC), and / or 1-cyclohexyl-2-morpholinoethylcarbodiimide (CMC); the N-hydroxy compound is selected from N-hydroxysuccinimide (NHS), sulfonyl... -N-hydroxysuccinimide (sulfonyl-NHS, Sulfo-NHS) and / or hydroxybenzotriazole (HOBt); or (b) a single-component triazine salt activator, such as 4-(4,6-dimethoxy-1,3,5-triazine-2-yl)-4-methylmorpholinium salt (DMTMM); or (c) a ureonium / phosphonium salt coupling agent, such as O-(7-azabenzotriazol-1-yl)-N,N,N',N'-tetramethylurea hexafluorophosphate (HATU), benzotriazol-1-yl-oxytripyrrolylphosphonium hexafluorophosphate (PyBOP) and / or (benzotriazol-1-yloxy)tris(dimethylamino)phosphonium hexafluorophosphate (BOP).

[0054] In some preferred embodiments, the preparation steps of the waterproof gel (adhesive) include:

[0055] 1. Prepare photoinitiator solution: Prepare a photoinitiator aqueous solution with a concentration of 0.1-0.4% (w / v), such as 0.15%, 0.2%, 0.25%, 0.3%, or 0.35%.

[0056] The photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP), sodium phenyl-2,4,6-trimethylbenzoylphosphonate (NAP), or a combination thereof. The present invention does not limit the solvent, but water or phosphate buffered saline (PBS) is preferred.

[0057] 2. Preparation of collagen matrix solution: Prepare a collagen matrix solution with a concentration of 0-20% (w / v). For example, dissolve collagen biomaterials in the above photoinitiator solution, and gently stir or intermittently shake continuously (e.g., for 30 minutes) until completely dissolved to obtain a clear and homogeneous collagen matrix solution.

[0058] The collagen material is gelatin or recombinant human collagen.

[0059] 3. Add monomer polymer: Prepare a prepolymer solution containing 2-20% (w / v) of methacrylamide derivative. For example, add methacrylamide derivative to the collagen matrix solution obtained in step 2, controlling the final concentration of methacrylamide derivative in the final mixed system to be 2-20% (w / v), and shake to mix to form a prepolymer solution with photopolymerization properties.

[0060] The methacrylamide derivative is selected from one or more of polyethylene glycol diacrylate (PEGDA), methacrylamide gelatin (GelMA), methacrylamide chitosan (CsMA), methacrylamide chondroitin sulfate (ChsMA), and methacrylamide dextran (DexMA).

[0061] Preferably, the process further includes a sterilization step: filtering the above-mentioned PG solution through a sterile filter membrane for sterilization. The sterilized solution is then stored at 4°C in the dark.

[0062] This invention does not limit the specific operations of solvent preparation, mixing, shaking, sterilization, filtration, and storage in the above steps; those skilled in the art can choose according to the actual situation. The embodiments described in this invention are some embodiments of this invention, but not all embodiments.

[0063] 4. Prepare coupling agent solution: Prepare a coupling agent solution with a concentration of 0.2-0.5% (w / v).

[0064] The coupling agent includes, but is not limited to, compounds selected from any one of the following or combinations thereof: (a) a combination system of carbodiimide and N-hydroxy compound, wherein the carbodiimide is selected from 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), N,N'-dicyclohexylcarbodiimide (DCC), N,N'-diisopropylcarbodiimide (DIC) and / or 1-cyclohexyl-2-morpholinoethylcarbodiimide (CMC); and the N-hydroxy compound is selected from N-hydroxysuccinimide (NHS), sulfonyl-N-hydroxysuccinimide (sulfonyl-NHS, Sulfo-NHS). (a) and / or hydroxybenzotriazole (HOBt); (b) single-component triazine salt activators, such as 4-(4,6-dimethoxy-1,3,5-triazine-2-yl)-4-methylmorpholinium salt (DMTMM); or (c) ureonium / phosphonium salt coupling agents, such as O-(7-azabenzotriazol-1-yl)-N,N,N',N'-tetramethylurea hexafluorophosphate (HATU), benzotriazol-1-yl-oxytripyrrolylphosphonium hexafluorophosphate (PyBOP), and / or (benzotriazol-1-yloxy)tri(dimethylamino)phosphonium hexafluorophosphate (BOP).

[0065] The present invention does not restrict the order of preparation of each solution in steps 1-4, and those skilled in the art can adjust it according to the actual situation.

[0066] 5. Preparation of the composite crosslinking solution: Mix the prepolymer solution and the coupling agent solution to obtain the composite crosslinking solution. Preferably, the main components of the composite crosslinking solution include: 0.1-0.4% (w / v) photoinitiator, no more than 20% (w / v) collagen material, 5-30% (v / v) methacrylamide derivative, and 0.05-0.5% (w / v) coupling agent;

[0067] Preferably, the mixing involves slowly adding a coupling agent solution dropwise to the prepolymer solution under continuous rapid mechanical stirring or vortex conditions. After the addition is complete, mixing continues (e.g., for 1-2 minutes) to ensure uniform dispersion.

[0068] Preferably, the process further includes a degassing step: removing air bubbles from the composite crosslinking solution to ensure the uniformity and optical transparency of the subsequently formed gel. For example, the composite crosslinking solution is transferred to a sterile centrifuge tube, placed in a preheated centrifuge, and centrifuged at 8000-12000 rpm for 1-2 minutes.

[0069] This invention does not limit the specific degassing operation; those skilled in the art can choose according to the actual situation. The embodiments described in this invention are some, but not all, of the embodiments.

[0070] 6. Constructing the gel framework: The composite cross-linking solution undergoes a cross-linking reaction to obtain a solid gel with a three-dimensional network structure.

[0071] Preferably, the reaction conditions are standing at 10-34°C.

[0072] The present invention does not limit the reaction time, for example, from 10 minutes to 2 hours.

[0073] This step has a significant impact on the dilution resistance and stability of the final waterproof gel, enabling it to more effectively resist the immediate flushing and dilution by bodily fluids such as tears, aqueous humor, and blood.

[0074] 7. High-temperature hydrolysis to regulate rheological properties: Heating the solid gel until its viscosity drops to 2000-20,000 mPa·s.

[0075] Preferably, the high temperature refers to a temperature not lower than 60°C. Preferably, the high temperature is 85-120°C, for example, 100±2°C.

[0076] Preferably, the heating method is a water bath; the heating time is 0.5 to 3 hours.

[0077] Preferably, in this step, the dynamic viscosity of the final product is precisely controlled within the range of 2000-20,000 mPa·s through real-time monitoring. Once the viscosity reaches the predetermined range, heating is immediately stopped, and the gel is cooled (e.g., at room temperature or 4°C) to obtain a waterproof gel with dilution resistance.

[0078] Preferably, steps 4-7 are carried out under light-protected conditions to prevent spontaneous polymerization of the prepolymer solution.

[0079] The waterproof gel prepared by this invention is colorless and transparent, meeting the requirements of high-definition vision in ophthalmic surgery. Specifically, the properties of the waterproof gel include, but are not limited to:

[0080] a) Optical performance: It has a transmittance of over 80% in the visible light band (400-700nm), and its refractive index (1.33-1.40) is similar to that of intraocular fluids such as aqueous humor and vitreous humor, as well as conventional irrigation fluids. When used intraocularly, it can minimize visual interference caused by interfacial light scattering and refractive difference.

[0081] b) Water resistance: The adhesive is thermodynamically insoluble in water or physiological fluids. Its mass loss rate after immersion in physiological saline at 37°C for 24 hours is less than 10%, demonstrating long-term structural stability and resistance to dissolution.

[0082] c) Dilution resistance: The covalent network within the adhesive effectively resists water dissolution and dilution. In dynamic fluid environments, it maintains its shape and structural integrity without significant disintegration or performance dilution, with a mass loss rate of less than 10% after 1 hour of underwater oscillation.

[0083] d) Erosion resistance: After the adhesive adheres to the ocular surface, it has the ability to resist tear erosion and effectively retain the tears. When tilted at 45° for 1 minute, the migration rate is less than 10%.

[0084] e) Viscoelasticity and Weissenberg Effect: The material exhibits typical linear viscoelastic behavior, and a normal stress difference, i.e., the Weissenberg effect, can be observed in rotational rheological tests (see Introduction to Colloid and Surface Chemistry, London Butterworths, 1966). This indicates that it has an elastic polymer network structure that can store and dissipate energy, which is beneficial for adhesion and resistance to transient deformation. The waterproof gel prepared in this invention has comprehensive adhesion properties including, but not limited to, shear strength of 45-55 kPa, tensile strength of 25-40 kPa, and interfacial toughness of 100-130 Jm. -2 .

[0085] f) Biocompatibility and Biodegradability: The main components of the gel exhibit good biocompatibility and biodegradability. In vitro cytotoxicity tests have confirmed that its extract has no significant toxicity to cells. Furthermore, the main components can be slowly degraded by enzymatic or hydrolytic processes, and the final products can be metabolized or excreted by the body.

[0086] g) Swelling property: It can reach swelling equilibrium in a body fluid environment, and its volume swelling ratio can be controlled between 1.0 and 5.0. In a near-low swelling state (e.g., 1.0-2.0), the material maintains structural stability and mechanical compatibility, adhering tightly to the tissue interface, making it suitable for sealing, repairing, and strengthening tissues. In a high swelling state (e.g., 2.0-5.0), the material can undergo moderate volume expansion, effectively filling irregular tissue defects or cavities.

[0087] h) Lubricity and tissue compatibility: The gel surface is rich in hydrophilic groups, which can form a hydration layer when in contact with tissue, significantly reducing the coefficient of friction and thus reducing mechanical damage between instruments or tissues during surgery.

[0088] i) Thixotropy: The material is in a high-viscosity gel state when at rest, and can remain at the target site; when subjected to shearing action such as injection, its viscosity drops rapidly (shear thinning), making it easy to deliver; after the shear force is removed, its viscosity can be quickly restored, ensuring that the material can be accurately filled during the operation and is not easily lost.

[0089] j) Osmotic pressure and biocompatibility: Through formulation control, the final gel's osmotic pressure can be precisely adjusted to a range isotonic with physiological fluids (approximately 280-320 mOsm / kg). This isotonic property avoids tissue and cell dehydration or edema caused by osmotic pressure differences, which is a necessary guarantee for its safe application in sensitive environments such as the intraocular region.

[0090] In summary, the waterproof gel prepared by this invention integrates key properties such as high transparency, water insolubility, erosion resistance, injectability, and biocompatibility. It possesses excellent resistance to water dilution and loss even before photocuring, allowing it to remain stably in the dynamic surgical environment filled with bodily fluids (such as aqueous humor, cerebrospinal fluid, and lymph), providing a reliable time window and morphological basis for subsequent photocuring operations. After light exposure, the material rapidly forms a robust interpenetrating network structure, achieving immediate and strong adhesion to moist tissues, thereby completing precise sealing, repair, or fixation functions. It is particularly suitable for ophthalmic surgeries with extremely high requirements for intraoperative visual field and postoperative optical quality: its high light transmittance and refractive index matching intraocular fluid mean that it hardly interferes with the surgeon's visual field when used on the ocular surface or inside the eye, and can minimize postoperative interface scattering, promoting visual function recovery; its inherent waterproof and erosion-resistant properties ensure that it can effectively fill wounds or fix tissues even in an eye with continuously flowing aqueous humor, and is not easily washed away. Therefore, the present invention can meet the requirements of underwater in-situ bonding and repair in ophthalmic surgeries (such as ocular surface tissue repair and bonding, retinal fixation).

[0091] This invention employs a two-step process—"low-temperature covalent cross-linking to construct the gel framework—high-temperature controllable hydrolysis to regulate rheological properties"—to create an in-situ curing system that combines good operability with excellent resistance to water dilution. This method is controllable, highly repeatable, and requires no complex equipment. The resulting gel exhibits stable properties, demonstrating good injectability and spreadability at room temperature, facilitating surgical procedures. In the moist ocular surface environment, its pre-formed covalent cross-linked network effectively resists immediate flushing and dilution by bodily fluids such as tears, aqueous humor, and blood, ensuring stable residence and precise positioning on the target tissue surface. Furthermore, under irradiation with a specific wavelength light source (such as 405nm or 365nm), a dense and robust secondary cross-linked network rapidly forms under moist conditions, achieving stable and reliable in-situ curing and tissue adhesion. In summary, this invention effectively solves the problems of traditional liquid adhesives being easily diluted, difficult to position, and having poor curing effects in moist surgical environments, providing a safe, efficient, and easy-to-operate sutureless wound closure solution for clinical surgery.

[0092] The technical solutions of the embodiments of the present invention will be clearly and completely described below with reference to the accompanying drawings. Obviously, the described embodiments are only some, not all, of the embodiments of the present invention. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort are within the scope of protection of the present invention.

[0093] While similar or equivalent steps, substances or materials, and reaction conditions may be used in the implementation of this invention, preferred steps, substances or materials, and reaction conditions are described herein.

[0094] When a range of values ​​is described herein, unless otherwise stated, the range is intended to include its endpoints and all integers and fractions within that range, and all values ​​within that range are capable of achieving the effects of the present invention.

[0095] Unless otherwise stated, all technical and scientific terms and abbreviations used herein have the meanings commonly understood by one of ordinary skill in the field of the invention or the field of application of such terms.

[0096] As used herein, the singular form of a word includes the plural, and vice versa. Therefore, “a,” “an,” and “the” generally include the plural form of the corresponding term. As used herein, “an embodiment” or “embodiment” refers to a specific feature, structure, or characteristic that may be included in at least one implementation of the invention. The phrase “in one embodiment” appearing in different places throughout this specification does not necessarily refer to the same embodiment, nor is it a single or selective embodiment that is mutually exclusive with other embodiments.

[0097] As used herein, “resistant,” “none,” “low,” or “reduced” means a state of resistance, reduction, or disappearance, or a reduction in a detectable or observable quantity. In some embodiments, one of the assessment tools described herein is used to measure this reduction or disappearance. In some embodiments, resistance or absence, or reduction or decrease, indicates a difference.

[0098] The rheological properties of waterproof gels refer to their ability to flow and deform under external forces (such as blinking, tear washing, and squeezing during drug administration).

[0099] Unless otherwise specified in the examples, the procedures should be performed under standard conditions or as recommended by the manufacturer. Reagents or instruments whose manufacturers are not specified are all commercially available or conventionally obtainable products.

[0100] Example 1: Preparation of Waterproof Adhesive (Gel)

[0101] 1. Preparation of photoinitiator solution: Take 5 mL of 0.25% (w / v) LAP (phenyl-2,4,6-trimethylbenzoylphosphonate lithium) solution (solvent is PBS) and add it to a 50 mL sterile centrifuge tube for later use.

[0102] 2. Introduction of polymerizable monomer (GelMA): Add 1.0 g of methacrylamide gelatin (GelMA, degree of substitution 90%). Place the mixture in a constant temperature water bath at 50±2℃ and gently shake for 30 minutes to ensure that the GelMA and collagen solution are fully and uniformly mixed to form a prepolymer solution. This step yields a GelMA concentration of 20% (w / v).

[0103] 3. Collagen matrix dissolution: Add 0.5g of collagen (Type I, derived from bovine Achilles tendon) to the centrifuge tube and place it in a constant temperature water bath at 40±2℃. Heat for 30 minutes, shaking for 10 seconds every 5 minutes until the collagen is completely dissolved, forming a homogeneous and clear mixed solution. This step yields a collagen concentration of 10% (w / v).

[0104] 4. Sterilization and Storage: Under safe red light conditions, sterilize the prepolymer solution obtained in step 3 by filtering it through a 0.22 μm polyethersulfone (PES) sterile filter membrane. Immediately after sterilization, transfer the solution to a 4°C refrigerator and store it in the dark. All subsequent operations must be performed in a dark environment.

[0105] 5. Preparation of coupling agent solution (prepare fresh before use): Under aseptic conditions, add 100 mg EDC / NHS powder to 500 μL of sterile PBS buffer (pH 7.4), vortex until completely dissolved, to obtain a 20% (w / v) EDC / NHS stock solution. Filter again through a 0.22 μm filter membrane for sterilization before use.

[0106] 6. Construction of the composite crosslinking system: Take approximately 5 mL of the prepolymer solution prepared in step 4 and transfer it to a sterile sample vial, placing it on a magnetic stirrer. While continuously stirring (approximately 200 rpm), slowly add the 20% EDC / NHS solution prepared in step 5 at a rate of 10 μL / s using a microinjection pump. After the addition is complete, continue stirring for 1 minute to ensure uniform mixing, obtaining a composite solution with a final EDC / NHS concentration of approximately 0.05% (w / v).

[0107] 7. Centrifugation and degassing: Transfer the above compound solution to sterile centrifuge tubes and place them in a centrifuge preheated to 40±2℃. Centrifuge at 10000 rpm for 1 minute to thoroughly remove dissolved and mixed air bubbles from the system.

[0108] 8. Filling and sealing: Fill a 5mL sterile syringe with the centrifuged clarified solution in a sterile operating table. Using a 1mL screw-type syringe, push the gel through the front end until no air bubbles remain. Then, reset the plunger of the 5mL syringe and tighten the sealing plug.

[0109] 9. Low-temperature covalent cross-linking (formation of the first network): Place the sealed syringe horizontally in a constant temperature (e.g., 10, 20, 30, 40°C) water bath or at room temperature, and let it stand for 10-30 minutes. During this stage, EDC / NHS activates collagen and GelMA molecules and forms a covalent cross-linked network, causing the system to gel and obtain a solid gel.

[0110] 10. High-Temperature Controlled Hydrolysis and Viscosity Adjustment (Rheological Property Optimization): The solid gel was placed in a water bath at 100±1℃ and boiled in boiling water for 1 hour. Subsequently, every 10 minutes, a small amount of gel sample was squeezed out through a syringe, and its apparent viscosity was measured using a rotational rheometer (using a cone-plate clamp, shear rate 1s⁻¹, 25℃). Heating was stopped when the measured viscosity value stabilized within the target range of 5000 mPa·s to 15000 mPa·s. The syringe was removed, and the sample was allowed to cool naturally to 25℃ at room temperature to obtain the waterproof adhesive (gel).

[0111] For ease of observation, a pigment was added to the gel in the examples. In actual preparation and application, the ophthalmic gel provided by this invention is colorless and transparent.

[0112] Example 2: Preparation of Waterproof Adhesive (Gel)

[0113] 1. Preparation of photoinitiator solution: Take 5 mL of 0.25% (w / v) LAP (phenyl-2,4,6-trimethylbenzoylphosphonate lithium) solution (solvent is PBS) and add it to a 50 mL sterile centrifuge tube for later use.

[0114] 2. Introduction of polymerizable monomer (GelMA): Add 0.5g of methacrylamide gelatin (GelMA, degree of substitution 90%). Place the mixture in a constant temperature water bath at 50±2℃ and gently shake for 30 minutes to ensure that the GelMA and collagen solution are fully and uniformly mixed to form a prepolymer solution. This step yields a GelMA concentration of 10% (w / v).

[0115] 3. Collagen matrix dissolution: Add 1.0 g of recombinant collagen to the centrifuge tube and place it in a constant temperature water bath at 40±2℃. Heat for 30 minutes, shaking for 10 seconds every 5 minutes until the collagen is completely dissolved, forming a homogeneous and clear mixed solution. This step yields a collagen concentration of 20% (w / v).

[0116] 4. Sterilization and Storage: Under safe red light conditions, sterilize the prepolymer solution obtained in step 3 by filtering it through a 0.22 μm polyethersulfone (PES) sterile filter membrane. Immediately after sterilization, transfer the solution to a 4°C refrigerator and store it in the dark. All subsequent operations must be performed in a dark environment.

[0117] 5. Preparation of coupling agent solution (prepare fresh for use): Under aseptic conditions, add 80 mg DMTMM powder (mass ratio 1:1) to 400 μL of sterile PBS buffer (pH 7.4), vortex until completely dissolved, to obtain a 20% (w / v) CMC / NHS stock solution, which should be sterilized by a 0.22 μm filter membrane before use.

[0118] 6. Construction of the composite crosslinking system: Take approximately 5 mL of the prepolymer solution prepared in step 4 and transfer it to a sterile sample vial, placing it on a magnetic stirrer. While continuously stirring (approximately 200 rpm), slowly add the 20% CMC / NHS solution prepared in step 5 at a rate of 10 μL / s using a microinjection pump. After the addition is complete, continue stirring for 1 minute to ensure uniform mixing, obtaining a composite solution with a final concentration of approximately 0.1% (w / v).

[0119] 7. Centrifugation and degassing: Transfer the above compound solution to sterile centrifuge tubes and place them in a centrifuge preheated to 40±2℃. Centrifuge at 10000 rpm for 1 minute to thoroughly remove dissolved and mixed air bubbles from the system.

[0120] 8. Filling and sealing: Fill a 5mL sterile syringe with the centrifuged clarified solution in a sterile operating table. Using a 1mL screw-type syringe, push the gel through the front end until no air bubbles remain. Then, reset the plunger of the 5mL syringe and tighten the sealing plug.

[0121] 9. Low-temperature covalent crosslinking (formation of the first network): Place the sealed syringe horizontally in a water bath at a constant temperature of 30°C or at room temperature, and let it stand for 20 minutes to obtain a solid gel.

[0122] 10. High-Temperature Controlled Hydrolysis and Viscosity Adjustment (Rheological Property Optimization): The solid gel was placed in a water bath at 60±1℃ for 1 hour. Subsequently, every 10 minutes, a small amount of gel sample was extruded through a syringe, and its apparent viscosity was measured using a rotational rheometer (using a cone-plate clamp, shear rate 1s⁻¹, 25℃). Heating was stopped when the measured viscosity value stabilized within the target range of 5000 mPa·s to 12000 mPa·s. The syringe was removed, and the gel was allowed to cool naturally to 25℃ at room temperature to obtain the waterproof gel.

[0123] Example 3: Preparation of Waterproof Adhesive

[0124] 1. Preparation of photoinitiator solution: Take 10 mL of 0.3% (w / v) LAP solution (solvent is PBS, pH 7.4) and add it to a 50 mL sterile centrifuge tube for later use.

[0125] 2. Introduction of polymerizable monomers: Add 2.0g of methacrylamide gelatin (GelMA, degree of substitution 60%) to the above centrifuge tube, place it in a constant temperature water bath at 50±2℃ and gently shake for 40 minutes until completely dissolved to obtain a prepolymer solution with a final GelMA concentration of 20% (w / v).

[0126] 3. Sterilization and storage: Sterilize the prepolymer solution by filtering it through a 0.22μm sterile filter membrane, and immediately transfer it to a 4℃ refrigerator for storage in the dark. All subsequent operations should be carried out in the dark.

[0127] 4. Preparation of coupling agent solution (prepare fresh for use): Under aseptic conditions, add 80 mg CMC / NHS powder (mass ratio 1:1) to 400 μL of sterile PBS buffer (pH 7.4), vortex until completely dissolved, to obtain a 20% (w / v) CMC / NHS stock solution, which should be sterilized again by passing through a 0.22 μm filter membrane before use.

[0128] 5. Construction of the composite crosslinking system: Take the prepolymer solution (about 8 mL) prepared in step 3 and transfer it to a sterile sample bottle. Stir continuously with a magnetic stirrer at 200 rpm, and slowly add the CMC / NHS stock solution from step 4. After the addition is complete, continue stirring for 2 minutes to obtain a composite solution with a final CMC / NHS concentration of 0.4% (w / v).

[0129] 6. Centrifugation and degassing: Transfer the composite solution to a sterile centrifuge tube, place it in a preheated centrifuge at 40±2℃, and centrifuge at 10000rpm for 1.5 minutes to remove air bubbles from the system.

[0130] 7. Filling and sealing: In a sterile operating table, fill a 5mL sterile syringe with the clear solution, and then use a 1mL screw-top syringe to push the solution until there are no air bubbles. Then, reset the plunger and tighten the sealing plug.

[0131] 8. Low-temperature covalent crosslinking: Place the sealed syringe horizontally at room temperature (25°C) for 10 minutes to form a covalent crosslinking network. The system then gels to obtain a solid gel.

[0132] 9. High-temperature controllable hydrolysis and viscosity adjustment: The solid gel was placed in a water bath at 120±1℃ and boiled for 0.5 hours. After that, a small amount of sample was squeezed out every 10 minutes and the apparent viscosity was measured with a rotational rheometer. When the viscosity stabilized at 6000~14000mPa・s, the heating was stopped and the gel was allowed to cool naturally to 25℃ at room temperature to obtain the waterproof gel.

[0133] Example 4: Preparation of Waterproof Adhesive

[0134] 1. Preparation of photoinitiator solution: Take 20 mL of 0.1% (w / v) LAP solution (solvent is PBS, pH 7.4) and add it to a 100 mL sterile centrifuge tube for later use.

[0135] 2. Dissolving the gelatin matrix: Add 4.0g of gelatin to the centrifuge tube and place it in a constant temperature water bath at 45±2℃ and stir for 30 minutes until the gelatin is completely dissolved to form a clear gelatin matrix solution.

[0136] 3. Introduction of polymerizable monomers: While stirring continuously, slowly add 4.0 mL of PEGDA (400) to the gelatin matrix solution, and continue stirring for 20 minutes until the mixture is homogeneous to form a prepolymer solution.

[0137] 4. Sterilization and storage: Under safe red light conditions, the prepolymer solution is sterilized by filtration through a 0.22μm sterile filter membrane and stored in a refrigerator at 4℃ in the dark. All subsequent operations should be conducted in the dark.

[0138] 5. Preparation of coupling agent solution: Under aseptic conditions, add 100 mg DMTMM powder to 500 μL of sterile PBS buffer (pH 7.4), vortex until completely dissolved, to obtain a 20% (w / v) DMTMM stock solution, which is sterilized by a 0.22 μm filter membrane before use.

[0139] 6. Construction of the composite crosslinking system: Take the prepolymer solution (about 10 mL) prepared in step 4 and transfer it to a sterile sample bottle. Stir continuously with a magnetic stirrer at 250 rpm. Add DMTMM stock solution dropwise at a rate of 10 μL / s. Continue stirring for 1.5 minutes after the addition to obtain a composite solution with a final DMTMM concentration of 0.5% (w / v).

[0140] 7. Centrifugation to remove air bubbles: Transfer the compound solution to a sterile centrifuge tube and centrifuge at 11,000 rpm for 1 minute at 40±2℃ to remove air bubbles.

[0141] 8. Filling and sealing: Fill a 5mL sterile syringe with the solution in the aseptic workbench, push it until there are no air bubbles, and then tighten the sealing plug.

[0142] 9. Low-temperature covalent crosslinking: Place the sealed syringe in a 30°C constant temperature water bath and let it stand for 20 minutes to gel the system and obtain a solid gel.

[0143] 10. High-temperature controllable hydrolysis and viscosity adjustment: The solid gel is placed in a water bath at 80±1℃ and heated for 3 hours. A small amount of sample is squeezed out every 10 minutes. The apparent viscosity is measured with a rotational rheometer. When the viscosity stabilizes at 7000~13000mPa・s, heating is stopped. The sample is cooled to 25℃ at room temperature to obtain the waterproof adhesive.

[0144] Example 5: Preparation of Waterproof Adhesive

[0145] 1. Preparation of photoinitiator solution: Take 10 mL of 0.4% (w / v) NAP solution (solvent is PBS, pH 7.4) and add it to a 50 mL sterile centrifuge tube for later use.

[0146] 2. Dissolving the gelatin matrix: Add 2.0g of gelatin to the centrifuge tube and place it in a constant temperature water bath at 42±2℃ and stir for 35 minutes until the gelatin is completely dissolved to form a homogeneous matrix solution.

[0147] 3. Introduction of polymerizable monomers: CsMA and ChsMA are added to the matrix solution in sequence and stirred for 25 minutes until they are mixed evenly to form a prepolymer solution (CsMA + ChsMA final concentration 5% (v / v)).

[0148] 4. Sterilization and storage: Under safe red light conditions, the prepolymer solution is sterilized by filtration through a 0.22μm filter membrane and stored in a refrigerator at 4℃ away from light. All subsequent operations should be conducted in the dark.

[0149] 5. Preparation of coupling agent solution: Under aseptic conditions, add 50 mg of PyBOP powder to 500 μL of sterile PBS buffer (pH 7.4), vortex until completely dissolved, to obtain a 10% (w / v) PyBOP stock solution.

[0150] 6. Construction of the composite crosslinking system: Take the prepolymer solution (about 8 mL) prepared in step 4 and transfer it to a sterile sample bottle. Stir continuously with a magnetic stirrer at 200 rpm. Add PyBOP stock solution dropwise at a rate of 9 μL / s. Continue stirring for 2 minutes after the addition to obtain a composite solution with a final PyBOP concentration of 0.5% (w / v).

[0151] 7. Centrifugation and degassing: Transfer the compound solution to a sterile centrifuge tube and centrifuge at 10,000 rpm for 2 minutes at 40±2℃ to completely remove air bubbles.

[0152] 8. Filling and sealing: Fill a 5mL sterile syringe with the clear solution in the aseptic workbench, push until there are no air bubbles, and then tighten the sealing plug.

[0153] 9. Low-temperature covalent crosslinking: Place the sealed syringe in a 20°C constant temperature water bath for 30 minutes. PyBOP catalyzes the crosslinking of carboxyl and amino groups between gelatin, CsMA, and ChsMA molecules to form a covalent network. The system then gels to obtain a solid gel.

[0154] 10. High-temperature controllable hydrolysis and viscosity adjustment: The solid gel was placed in a water bath at 100±1℃ and boiled for 1 hour. A small amount of sample was squeezed out every 10 minutes. The apparent viscosity was measured using a rotational rheometer (cone plate clamp, shear rate 1s⁻¹, 25℃). When the viscosity stabilized at 5000-15000mPa・s, heating was stopped and the sample was allowed to cool naturally to 25℃ at room temperature to obtain the waterproof adhesive.

[0155] To clearly demonstrate the formulation components, concentrations, and core design advantages of each embodiment of the present invention, the key parameters of Examples 1-5 are summarized in the table below, each with its own performance emphasis, and can be adapted to different clinical application scenarios:

[0156] Example 1 2 3 4 5 Photoinitiator (w / v) LAP (0.25%) LAP (0.25%) LAP (0.3%) LAP (0.1%) NAP (0.4%) Collagen-based materials (w / v) Collagen (10%) Collagen (20%) 0% Gelatin (20%) Gelatin (10%) Methacrylamide derivatives (v / v) GelMA (20%) GelMA (10%) GelMA (20%) PEGDA (6%) CsMA + ChsMA (5%) Coupling agent (w / v) EDC / NHS (0.05%) DMTMM (0.1%) CMC / NHS (0.4%) DMTMM (0.5%) PyBOP (0.5%) Viscosity range (mPa·s) 5000-15000 5000-12000 6000-14000 7000-13000 5000-15000

[0157] Test Example 1: Waterproof / Water-resistant Dilution Performance Test

[0158] To evaluate the anti-dilution and anti-erosion capabilities of the waterproof adhesive (i.e., the uncured intermediate product) prepared by the present invention in a dynamic aqueous environment, this experiment compared the behavior differences between the prepolymer obtained by the method of the present invention and two control materials in different states in simulated body fluids.

[0159] Sample Preparation: Example 1 Sample: Waterproof adhesive prepared according to the method of Example 1. This sample represents the fluid form of the invention prior to surgical injection.

[0160] Control Group 1: A pristine liquid prepolymer solution without any covalent crosslinking or high-temperature hydrolysis treatment (preparation method is the same as steps 1 to 4 of Example 1). This sample represents the initial state of most low-viscosity, injectable ocular hydrogel materials that currently rely on in-situ photopolymerization.

[0161] Control group 2: The preparation method is the same as steps 1 to 9 of Example 1, that is, using a solid gel (prepolymer) that has been cross-linked but has not undergone high-temperature controlled hydrolysis.

[0162] The experimental method is as follows: 1 mL of the waterproof adhesive from Example 1 and the control sample were placed in 20 mL of PBS buffer (pH 7.4, simulating physiological body fluid environment) and shaken at 50 rpm in a constant temperature shaker at 37°C. The morphological changes of the samples were observed after 1 minute. The results showed that: Control sample 1 (liquid prepolymer solution): rapidly diffused and dissolved within 1 minute after the start of shaking, completely mixed with the buffer, and lost its original morphology and positioning ability. Control sample 2 (solid gel): maintained its complete solid morphology throughout the entire experimental period, without any visible dissolution or dispersion, confirming the insoluble properties imparted by the covalent cross-linked framework. However, its rigid structure also resulted in its lack of deformation ability, making it unsuitable for injection application. Therefore, it was difficult to achieve effective adaptation and positioning on dynamic and irregular wound surfaces, greatly limiting its clinical usability. Sample 1 (the adhesive of this invention): under the same conditions, its morphology remained intact, with no obvious dissolution or dispersion, exhibiting excellent water dilution stability comparable to control sample 2. Simultaneously, it exhibited certain viscoelastic deformation in the shaking environment, unlike the rigid solid of control sample 2.

[0163] Experimental results are as follows Figure 1As shown, this invention, through a process of "first constructing a covalently cross-linked framework," endows the material with water-insoluble properties. The subsequent "controlled hydrolysis" step, while maintaining this insolubility, imparts the necessary deformability and viscoelasticity to the material. In contrast, conventional liquid prepolymers rapidly fail in dynamic aqueous environments due to the lack of a stable network framework. The erosion and dilution resistance exhibited by the gel of this invention is of crucial practical significance for maintaining material stability and ensuring effective adhesion and repair in ophthalmic surgical environments where aqueous humor and tear film flow are constantly present.

[0164] Test Example 2: Gel Positioning Retention Ability Test

[0165] To verify the effect of the two-step process of "low-temperature covalent crosslinking-high-temperature controllable hydrolysis" described in this invention on improving the anti-flow properties of the gel, an inclined plate test simulating the curvature of the eyeball was designed in this experiment. Three control group samples were set up, and their preparation process was basically the same as that in Example 1, but the key low-temperature crosslinking and high-temperature hydrolysis control steps were omitted, thus obtaining a low-viscosity liquid prepolymer solution containing only photopolymerizable monomers and photoinitiators without rheological control. This state can represent most of the current ocular hydrogel materials that rely on photocuring.

[0166] The experimental method is as follows: Under simulated body surface temperature (37°C), 100 μL of the waterproof adhesive prepared in Example 1 and control group 3 samples were respectively dropped vertically onto a horizontal plate marked with fine graduations. The plate was quickly tilted to a 45° angle, left to stand for 1 minute, and then the sample morphology was photographed and recorded. The migration rate of the sample front was calculated. Results ( Figure 2 The results show that the adhesive prepared in Example 1 has a migration rate of only 6.2%, indicating that it hardly flows on inclined surfaces and has good positional stability. In contrast, the migration rate of control group 3 is as high as 107.6%, with significant flow and spreading, proving that it cannot maintain its shape and position in non-planar environments. These results demonstrate that the present invention can endow materials with excellent in-situ retention and anti-flow properties in dynamic humid environments by precisely controlling the viscoelastic properties of the gel. Untreated conventional liquid materials, due to their high fluidity, easily flow away from the target area and cannot stably remain on curved surfaces, which is a key limitation restricting the practical application of many current hydrogel adhesives in clinical practice. Therefore, the gel prepared by the present invention can effectively solve the technical problems of difficult intraoperative adhesive positioning, unnecessary adhesion, and complications.

[0167] Test Example 3: Determination of Adhesion Strength of Ex vivo Corneal Tissue

[0168] To objectively evaluate the overall adhesion performance of the waterproof adhesive described in this invention, this comparative example was conducted on an in vitro porcine cornea model, and its mechanical properties were systematically compared with those of two commonly used commercial adhesives in clinical settings—cyanoacrylate super glue and medical fibrin glue—according to the standardized test methods of the American Society for Testing and Materials (ASTM). The toughness, shear resistance, and tensile strength of the adhesive interface were quantitatively evaluated through 180-degree peel, shear, and tensile tests.

[0169] The experiment used fresh porcine corneas to prepare a standardized anterior lamellar defect model, which was sealed using the adhesive prepared in Example 1 of this invention, cyanoacrylate glue, and fibrin glue, respectively. After bonding, the corneal tissue was processed into test samples of specific sizes, and polymethyl methacrylate (PMMA) backing plates were fixed to both ends of the samples to enhance clamping stability. Subsequently, various mechanical tests were performed on a dynamic mechanical analyzer (Q800, TA Instruments) at a uniform rate of 5 mm / min. Each adhesive was tested in triplicate to ensure the repeatability and statistical reliability of the test data.

[0170] Test results ( Figure 3 The results show that the waterproof gel prepared by this invention exhibits significantly better overall adhesive performance than the control group. Specifically, the adhesive interface toughness of Example 1 is higher than 120 Jm⁻², demonstrating excellent peel resistance and energy absorption capacity of the adhesive interface; the shear strength and tensile strength reach 51.7±2.3 kPa and 30.7±2.1 kPa, respectively, indicating that the gel has good shear and tensile anchoring ability on the corneal surface. In contrast, the strength indicators of the fibrin adhesive are significantly lower than those of the gel of this invention. Although cyanoacrylate adhesives exhibit high initial adhesion strength in dry environments, they are prone to adhesion failure on moist tissue surfaces and have problems such as curing exothermics and tissue irritation, which limit their application in ophthalmic surgery.

[0171] Test Example 4: In vitro biocompatibility evaluation

[0172] To evaluate the cell compatibility of the waterproof gel of this invention, this embodiment uses human corneal epithelial cells (HCEC) and human corneal stromal cells (HCSC) in an indirect co-culture system to systematically evaluate the in vitro biocompatibility of the material prepared in Example 1. First, HCEC and HCSC were cultured at 1×10⁻⁶ cells per cell line. 4Cells were seeded at a density of [number] cells / well in 24-well plates and cultured at 37°C and 5% CO2 for 24 hours to allow for full cell adhesion. The gel precursor solution prepared in Example 1 was placed in a mold with a diameter of 6 mm and a thickness of approximately 1 mm and cured under 10 mW / cm² 405 nm blue light for 60 seconds to prepare homogeneous gel discs. These gel discs were placed on the upper layer of a Transwell chamber and indirectly co-cultured with HCECs or HCSCs in the lower layer for 48 hours.

[0173] Cells co-cultured were stained using a live / dead cell fluorescence double staining kit, and the results were observed under a fluorescence microscope. Figure 4 The results showed that both cell types exhibited good spreading morphology after co-culturing with the gel, with continuous and intact cell layers, clear boundaries, and no obvious aggregation of PI red fluorescence (dead cells) (see Appendix). Figure 4 (A and B). Cell viability was further assessed using the CCK-8 assay. The results showed that the relative viability of HCECs and HCSCs co-cultured with the gel were (97.6±0.3)% and (96.8±0.4)%, respectively, with no statistically significant difference compared to the blank control group (no material) (p>0.05) (see Appendix). Figure 4 (C). The above results indicate that the waterproof gel prepared in this invention exhibits good biocompatibility with both major cell types derived from human cornea under in vitro conditions, and does not show significant cytotoxicity.

[0174] Test Example 5: Waterproof Adhesive Performance Test

[0175] The waterproof adhesives prepared in Examples 1-4 were tested for dilution resistance, positioning retention ability, adhesive strength, optical properties, biocompatibility, and water absorption. The test methods were the same as those for the corresponding performance tests in Example 1. The specific test results are shown in the table below.

[0176] Test metrics Example 1 Example 2 Example 3 Example 4 Example 5 Dilution resistance - mass loss rate after 1 hour of underwater oscillation (%) 7.2±0.5 5.5±0.3 6.3±0.4 5.7±0.3 6.8±0.4 Positioning retention capability -45° tilt for 1 minute, migration rate (%) 6.2±0.6 4.2±0.2 5.8±0.5 4.9±0.4 5.5±0.5 Adhesive strength - shear strength (kPa) 51.7±2.3 47.6±3.1 48.9±2.1 53.2±1.8 50.5±2.0 Bond strength - tensile strength (kPa) 30.7±2.1 35.5±8.5 29.5±1.9 32.1±1.9 31.2±2.0 Bond strength - interfacial toughness (Jm-2) ≥120 ≥125 ≥115 ≥125 ≥122 Optical performance - 550nm transmittance (%) ≥90 ≥90 ≥92 ≥91 ≥90 Optical properties - refractive index 1.35±0.01 1.34±0.01 1.35±0.01 1.36±0.01 1.37±0.01 Biocompatibility - Cell viability (%) 97.2±0.4 98.5±0.2 98.1±0.2 97.5±0.3 97.2±0.3 Swelling properties - water absorption ratio 1.3±0.1 1.1±0.1 1.2±0.1 4.5±0.1 1.2±0.1 Photocurability - Curing time (s) 60 50 55 40 50

[0177] Test Example 6: Underwater Tissue Adhesion Experiment

[0178] Two pieces of porcine small intestinal mucosa tissue were placed in a culture dish containing physiological saline (simulating a body fluid environment), keeping the tissues moist underwater throughout the process. Using a microsyringe, approximately 50 μL of the waterproof gel prepared in Example 1 was applied to the surface of one of the small intestinal mucosa pieces. The other small intestinal mucosa piece was then placed next to the gel-coated tissue piece, and the area was subsequently irradiated with a 405 nm blue light source (light power density 15 mW / cm²) for 30 seconds to complete the photocuring. Figure 5 (A). Experimental results: After light curing, the two pieces of small intestinal mucosa tissue were tightly bonded together by the waterproof gel; when one piece of tissue was lifted by its edge using tweezers, both pieces of tissue could be lifted simultaneously without separation. Figure 5(B) Even when the tweezers are shaken, the adhesive interface remains intact. Experiments demonstrate that this waterproof gel can effectively achieve tissue adhesion underwater, and the adhesive strength after light curing can stably support the weight of the tissue itself, proving that it has reliable tissue adhesion performance and erosion resistance in humid / underwater surgical environments, and can meet the tissue repair needs of clinical surgery.

[0179] Animal experiments: Experiments on corneal defect repair in animals

[0180] To systematically verify the tissue repair efficacy of the gel of this invention in a real physiological environment, this embodiment conducts in vivo functional evaluation by establishing a rabbit corneal lamellar defect model. Twelve healthy New Zealand white rabbits were randomly divided into an experimental group and a control group, with six rabbits in each group. A standard lamellar defect with a diameter of 3 mm and a depth of 200 μm was created in the center of the cornea. The experimental group used the gel prepared by the method in Example 1 for defect closure; the control group used a composite solution that had only completed steps 1-4 and had not undergone cross-linking treatment as a control material. Before photocuring, both groups of wounds were rinsed with physiological saline to simulate real clinical interference conditions such as intraoperative bleeding, aqueous humor exudation, and tear flushing. Subsequently, in-situ curing was completed by irradiation with a light power density of 10 mW / cm² for 60 seconds.

[0181] Immediate postoperative OCT imaging ( Figure 6 The results showed that in the experimental group, the corneal defect was completely covered by gel, with the material forming a continuous and consistent curved structure with the corneal tissue; in the control group, the gel coverage was incomplete, with material remaining only at the edge of the wound. On the 7th postoperative day, a systematic evaluation using slit-lamp examination, sodium fluorescein staining, and anterior segment OCT revealed that in the experimental group, the gel remained firmly adhered to the wound without displacement or detachment, the corneal epithelium healed completely within 7 days, tissue transparency was good, and no obvious inflammation or neovascularization was observed; the control group, however, showed incomplete gel coverage, delayed wound closure, and significant signs of fibrosis.

[0182] This experiment demonstrates that the gel prepared by the controlled cross-linking process of this invention maintains a stable three-dimensional structure and interfacial adhesion in a humid environment simulating real surgical disturbances, effectively supporting and promoting corneal tissue regeneration and repair. In contrast, conventional liquid hydrogels are easily affected by fluid erosion under the same conditions, leading to material loss, incomplete coverage, and reduced repair effectiveness. These results further verify that the material of this invention possesses excellent anti-dilution properties, morphological stability, and biofunctionality in complex ocular surface environments.

[0183] Animal experiment: Seamless conjunctival transplantation experiment in rabbits

[0184] To verify the clinical application potential of the gel of this invention for graft adhesion, this embodiment conducted a comparative study on unilateral conjunctival transplantation in a rabbit eye model. Eight healthy New Zealand white rabbits were selected to establish an autologous conjunctival transplantation model and randomly divided into an experimental group and a control group. The experimental group used the gel prepared in Example 1 for unilateral adhesion and fixation of the graft; the control group used conventional 10-0 nylon sutures for interrupted suturing. Postoperative observation lasted for 14 days, and graft survival, conjunctival healing progress, and the occurrence of complications were systematically evaluated.

[0185] Postoperative observation results ( Figure 7 The results showed that all grafts in the experimental group survived well, with no graft detachment, displacement, or hematoma formation. Conjunctival wound edges began to heal on postoperative day 3, and the transplanted area healed well and showed tight tissue integration on day 7. During the 14-day observation period, no complications such as infection, bleeding, or foreign body reaction occurred. The healing effect was essentially equivalent to that of the traditional suture group, achieving similar surgical repair goals.

[0186] This experiment demonstrates that the gel of this invention can serve as a safe and effective bio-adhesive, successfully achieving seamless fixation of conjunctival grafts. Compared with traditional suturing techniques, this gel bonding method not only simplifies the surgical procedure and shortens the operation time, but also reduces the risk of tissue trauma, foreign body reaction, and complications caused by sutures. Simultaneously, it promotes rapid healing and functional recovery of the transplanted tissue, showing promising prospects for clinical translation.

[0187] Clinical trial: Conjunctival graft adhesion after pterygium excision

[0188] Pterygium is a common proliferative disease of the ocular surface with a high global incidence. Its standard treatment involves surgical excision combined with autologous conjunctival transplantation. The core step is to securely fix the free conjunctival graft to the scleral recipient bed. Traditional techniques rely entirely on surgical sutures for fixation. Existing liquid tissue adhesives, due to the continuous tear washing, tissue bleeding, and moist environment of the ocular surface, are easily diluted, washed away, or displaced before curing, making it difficult to form a reliable bond. Therefore, they cannot replace sutures in such delicate, moist, and dynamic microsurgeries.

[0189] To evaluate the practical application effect of the gel of this invention in complex ocular surface surgery, this embodiment conducted a prospective clinical verification in two cases of pterygium excision combined with conjunctival transplantation. The surgery was performed under peribulbar anesthesia. After complete excision of the pterygium lesion, an autologous conjunctival transplant flap of appropriate size was prepared. During the operation, the filler prepared in Example 1 of this invention was used as a biological adhesive to replace traditional sutures for bonding and fixing the transplant flap. The prepolymer was uniformly injected into the gap between the transplant flap and the recipient bed. Excess material at the interface was gently pressed out and residual gel was removed. Finally, in-situ curing was completed by irradiation with a curing light source at a wavelength of 365 nm and an intensity of 18 mW / cm² for 60 seconds.

[0190] Postoperative observation and follow-up results ( Figure 8 The results showed that the gel-bonded graft adhered firmly to the recipient tissue, without displacement or detachment despite postoperative tear flushing and eyelid movements. After healing, the conjunctival surface was smooth, with good tissue integration, no granulation tissue hyperplasia or scar formation, and the limbal anatomy recovered well. This clinical validation demonstrates that the present invention can meet the seamless bonding requirements of conjunctival transplantation in pterygium surgery. Its suitable operating viscosity and photo-triggered curing characteristics provide ample operational window for precise intraoperative positioning, and the stable adhesive interface formed after curing effectively resists the effects of the ocular moist environment and physiological mechanical stress.

[0191] This technology not only simplifies surgical procedures and shortens surgical time, but also has the potential to significantly reduce the risk of complications such as inflammatory reactions, foreign body sensation, and scar hyperplasia caused by suture irritation. It demonstrates good biocompatibility, operational safety, and clinical applicability, and has significant clinical translational value and promising prospects for widespread application.

[0192] The specific embodiments described above do not constitute a limitation on the scope of protection of this invention. Those skilled in the art should understand that various modifications, combinations, sub-combinations, and substitutions can be made based on the inventive principles and other factors. Any modifications, equivalent substitutions, and improvements made within the spirit and principles of this invention should be included within the scope of protection of this invention.

Claims

1. A method for preparing a medical waterproof adhesive, characterized in that, Includes the following steps: A solid gel with a three-dimensional cross-linked network connected by amide bonds was obtained by using a coupling agent to catalyze the condensation reaction of collagen materials and / or methacrylamide derivative molecules. The solid gel is placed in an aqueous environment to partially break the amide bonds inside the solid gel until the solid gel becomes a viscoelastic fluid, thus obtaining a viscoelastic waterproof adhesive.

2. The preparation method according to claim 1, characterized in that, Includes the following steps: Step 1) Prepare a composite crosslinking solution, the components of which include: 0.1-0.4% (w / v) photoinitiator, no more than 20% (w / v) collagen material, 5-30% (v / v) methacrylamide derivative and 0.05-0.5% (w / v) coupling agent; Step 2) The composite cross-linking solution is placed at the reaction temperature to react and obtain a solid gel with a three-dimensional cross-linked network with amide bonds as the connection points; Step 3) The solid gel is placed in a high-temperature aqueous environment for hydrolysis. When the viscosity drops to 2000-20000 mPa·s, a viscoelastic waterproof adhesive is obtained; the high temperature refers to 60-150°C.

3. The preparation method according to claim 1, characterized in that, The collagen material is gelatin or recombinant human collagen. The coupling agent is a combination of a carbodiimide and an N-hydroxy compound, wherein the carbodiimide is selected from one or more of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide, N,N'-dicyclohexylcarbodiimide, N,N'-diisopropylcarbodiimide, and 1-cyclohexyl-2-morpholinoethylcarbodiimide; and the N-hydroxy compound is selected from one or more of N-hydroxysuccinimide and sulfonyl-N-hydroxysuccinimide. Alternatively, the coupling agent may be a triazineonium salt coupling agent; Alternatively, the coupling agent may be a ureonium / phosphoniumium salt coupling agent selected from one or more of the following: O-(7-azabenzotriazol-1-yl)-N,N,N',N'-tetramethylurea hexafluorophosphate, benzotriazol-1-yl-oxytripyrrolidinylphosphine hexafluorophosphate, or (benzotriazol-1-yloxy)tris(dimethylamino)phosphonium hexafluorophosphate. The methacrylamide derivative is selected from one or more of polyethylene glycol diacrylate (PEGDA), methacrylamide gelatin (GelMA), methacrylamide chitosan (CsMA), methacrylamide chondroitin sulfate (ChsMA), and methacrylamide dextran (DexMA).

4. The preparation method according to claim 2, characterized in that, The photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonate (LAP), sodium phenyl-2,4,6-trimethylbenzoylphosphonate (NAP), or a combination thereof; the reaction temperature refers to standing at 10-34°C; the high temperature refers to 85-120°C.

5. The waterproof adhesive prepared by any one of the preparation methods described in claims 1-4.

6. The waterproof adhesive as described in claim 5, characterized in that, Under light irradiation, the waterproof adhesive forms a hydrophilic polymer network on a three-dimensional cross-linked network, thus becoming a solid waterproof adhesive.

7. The waterproof adhesive as described in claim 5, characterized in that, The waterproof adhesive has waterproof, dilution-resistant, and erosion-resistant properties.

8. The waterproof adhesive as described in claim 5, characterized in that, Light transmittance ≥ 80%; refractive index 1.33-1.40; volume swelling ratio 1.0-5.0; mass loss rate less than 10% after 1 hour of underwater oscillation; migration rate less than 10% after 1 minute of 45° tilting; shear strength 45-55 kPa; tensile strength 25-40 kPa; interfacial toughness 100-130 Jm -2 .

9. The application of the waterproof adhesive as described in claim 7 in the preparation of medical devices for use in humid environments.

10. The use of the waterproof adhesive as described in claim 7 in the preparation of a medical device for ophthalmic surgery.

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