Injectable double-layer drug release hydrogel and preparation method and application thereof

Abstract

The application discloses an injectable double-layer drug sustained-release hydrogel and a preparation method and application thereof, and relates to the field of drug sustained-release hydrogels. The swelling and degradation curve and mechanical property detection of the hydrogel show that, when the swelling rate ratio of the inner layer hydrogel to the outer layer hydrogel is 14% to 10%, the secondary damage of the optic nerve caused by extrusion can be avoided; meanwhile, the hydrogel has suitable mechanical strength and adhesion to cope with the in-vivo environment. The hydrogel precursor solution has good injectability, and the gelation time can be controlled within 2 minutes, which perfectly matches the short time window of the optic nerve decompression surgery, and has high clinical practical value. The application solves the major surgical risk of cerebrospinal fluid leakage when the optic nerve is locally administered, and solves the problems of the traditional local drug carrier, such as lack of in-situ stability and easy displacement. The application can synergistically regulate the injury microenvironment and promote the long-acting sustained-release of nerve regeneration, successfully integrates multiple functions such as surgical safety, drug long-acting sustained-release, biocompatibility and operation convenience, and provides an unprecedented effective strategy for solving the clinical problem of the optic nerve injury.

Description

Technical Field

[0001] This invention relates to the field of biomedical materials and controlled drug release technology, specifically to an injectable bilayer drug sustained-release hydrogel, its preparation method, and its application. Background Technology

[0002] Optic nerve injury is a major blinding eye disease leading to irreversible vision loss, such as glaucoma and traumatic optic neuropathy. Its core pathological mechanism involves apoptosis of retinal ganglion cells and axonal degeneration. Currently, there are no effective treatments to promote optic nerve repair and regeneration. The main challenge of drug therapy lies in the fact that the optic nerve, as part of the central nervous system, is strictly regulated by the blood-retinal and blood-brain barriers, making it difficult for most drugs to effectively reach the lesion site when administered systemically (orally or intravenously), resulting in extremely low bioavailability.

[0003] To overcome these barriers, previous studies have focused on local drug delivery strategies. Existing technologies can be mainly categorized as follows:

[0004] Intravitreal injection

[0005] Intravitreal injection is a common method for local drug administration in posterior segment eye diseases. This method involves directly injecting medication into the vitreous cavity, where it acts on the cell bodies of retinal ganglion cells. However, this method has an inherent limitation for treating optic nerve injuries: the drug needs to undergo a long retrograde transport from the posterior segment of the eye via axons to reach the site of optic nerve injury. After the injury occurs, axonal transport function is already severely impaired, preventing the drug from reaching the primary lesion in a timely and sufficient manner for effective intervention, thus resulting in minimal therapeutic effect.

[0006] For precise treatment, local drug delivery technology targeting the optic nerve is a promising direction, with direct local administration of drugs near the optic nerve being a key area. Existing research has explored the direct application of drugs or drug carriers to the surface of the optic nerve during optic nerve decompression surgery.

[0007] While the aforementioned local optic nerve drug delivery techniques offer the possibility of sustained drug release, their direct application to local optic nerve drug delivery presents the following serious drawbacks and unresolved technical challenges:

[0008] 1. Inability to effectively seal dural incisions poses a risk of cerebrospinal fluid (CSF) leakage: The optic nerve is encased in the dura mater, whose cavities communicate with the intracranial cavity and are filled with CSF. Any procedure that cuts the optic nerve sheath (dura mater) will lead to CSF ​​leakage, causing serious complications such as low intracranial pressure. Existing hydrogel technology primarily focuses on drug delivery and release; the material itself lacks sufficient bioadhesion or film-forming sealing properties, making it impossible to reliably seal dural incisions while administering medication. This is a key obstacle to translating laboratory concepts into safe clinical procedures.

[0009] 2. Lack of in-situ stability, easy displacement leading to poor efficacy and potential risks: There is tissue pulsation and cerebrospinal fluid flow around the optic nerve. Ordinary hydrogels are only filled by simple injection, and their adhesion to nerve tissue is weak, making them extremely prone to displacement, detachment, or being washed away. This not only prevents the drug from acting continuously at the lesion site, but the detached material may also pose a risk to surrounding tissues.

[0010] 3. Single materials struggle to simultaneously achieve occlusion, sustained release, and biocompatibility: Highly adhesive materials may cause toxicity or compressive damage to fragile nerve tissue; while biocompatible sustained-release materials often lack the immediate adhesion and occlusion capabilities required for surgery. Current technologies lack a multifunctional integrated design that can synergistically address "instantaneous occlusion," "long-term sustained release," and "neurocompatibility."

[0011] In summary, the existing technology lacks a local drug delivery system that can be used safely, effectively, and conveniently in optic nerve decompression surgery, achieving both long-acting drug release and reliable closure of the dura mater to prevent cerebrospinal fluid leakage.

[0012] The specific technical problems to be solved include:

[0013] 1. To address the problem that systemic and intravitreal drug delivery cannot effectively target optic nerve lesions. Due to the presence of the blood-eye / blood-brain barrier and axonal transport damage, the bioavailability of systemic drug delivery is extremely low, while intravitreal drug delivery has a lag effect, making it difficult for the drug to effectively reach the primary site of optic nerve injury.

[0014] 2. Addressing the significant surgical risk of cerebrospinal fluid leakage during local administration of medication to the optic nerve. After creating a dura mater window, such as in optic nerve canal decompression surgery, a barrier material is needed that can immediately and reliably seal the incision to prevent cerebrospinal fluid leakage.

[0015] 3. It solves the problems of traditional local drug delivery carriers lacking in-situ stability and being prone to displacement. The dynamic characteristics of the environment surrounding the optic nerve require the drug delivery carrier to adhere firmly to the tissue surface and remain there for a long time without being washed away or displaced by cerebrospinal fluid.

[0016] 4. To provide a long-acting sustained-release platform that can synergistically regulate the damaged microenvironment and promote nerve regeneration. Single drugs are insufficient to address the complex pathological processes following optic nerve injury (such as inflammation, oxidative stress, and myelin sheath disintegration), necessitating an intelligent carrier that can simultaneously deliver multiple functional factors and control their release behavior. Summary of the Invention

[0017] To address the technical deficiencies of existing technologies, this invention provides an injectable bilayer drug sustained-release hydrogel, its preparation method, and its application.

[0018] The technical solution adopted in this invention is: an injectable bilayer drug sustained-release hydrogel, wherein the hydrogel comprises an inner hydrogel and an outer hydrogel, the inner hydrogel is composed of methacrylamide gelatin (GelMA) and 2-mercapto-1-methylimidazole through photocrosslinking, and the inner hydrogel also loads and encapsulates drug nanoparticles, and the outer hydrogel is composed of methacrylamide gelatin (GelMA) and oxidized dextran (ODex) through photocrosslinking.

[0019] The concentrations of methacrylamide gelatin (GelMA) and 2-mercapto-1-methylimidazole in the inner hydrogel are 20% w / v and 5% w / v, respectively.

[0020] The drug-encapsulated nanoparticles are mesoporous polydopamine nanoparticles (MPDA@CLE) encapsulating chlormastine (CLE).

[0021] The concentrations of methacrylamide gelatin (GelMA) and oxidized dextran (ODex) in the outer hydrogel are 20% w / v and 5% w / v, respectively.

[0022] The photocrosslinking is performed by irradiating the LAP photoinitiator with blue-violet light at a wavelength of 405 nm.

[0023] A method for preparing an injectable bilayer drug-release hydrogel includes the following steps:

[0024] S1. Synthesis of mesoporous polydopamine nanoparticles (MPDA): Dopamine hydrochloride and triblock copolymer F127 were added to anhydrous ethanol and deionized water. The mixture was magnetically stirred at room temperature until the solid was completely dissolved. While continuously stirring, 1,3,5-trimethylbenzene (TMB) was added, and the mixture was ultrasonically emulsified at 35°C until the solution became a uniform milky white. The solution was then transferred to a 40°C water bath, and ammonia was added dropwise while maintaining stirring. After the addition was complete, the mixture was protected from light and the reaction continued until the reaction was finished. The supernatant was discarded by high-speed refrigerated centrifugation. The precipitate was washed with a mixture of anhydrous ethanol and deionized water, and then with a mixture of anhydrous ethanol and acetone, centrifuged after each wash. The final precipitate was freeze-dried to obtain black MPDA powder.

[0025] S2. Preparation of MPDA-loaded chlormastine (MPDA@CLE): MPDA powder was dispersed in chlormastine phosphate buffer solution and shaken at room temperature in the dark. After the reaction, the precipitate was collected by centrifugation, the precipitate was gently washed with PBS, and freeze-dried to obtain MPDA@CLE powder.

[0026] S3. Preparation of inner layer hydrogel solution: GelMA was dissolved in PBS to prepare a 20% solution. 5% 2-mercapto-1-methylimidazole and 0.25% photoinitiator LAP were added. The mixture was stirred until homogeneous and MPDA@CLE nanoparticles loaded with chlormastine were added. Photocrosslinking was performed to obtain an inner layer hydrogel solution loaded with MPDA@CLE nanoparticles.

[0027] S4. Preparation of outer hydrogel solution: Dissolve GelMA in PBS to prepare a 20% solution, add 5% ODex and 0.25% photoinitiator LAP, stir evenly, and photocrosslink to obtain the outer hydrogel solution.

[0028] In step S1, the volume ratios of anhydrous ethanol to deionized water and anhydrous ethanol to acetone are 1:1 and 2:1, respectively.

[0029] In step S2, the concentration of MPDA powder dispersed in chlormastine phosphate buffer solution is 1 mg / mL.

[0030] Application of an injectable bilayer drug-releasing hydrogel in the preparation of drugs for local administration of optic nerve injury.

[0031] The aforementioned drug for local administration of optic nerve injury is obtained by injecting the inner hydrogel solution of an injectable bilayer drug-releasing hydrogel into the target site within the optic nerve sheath, irradiating it with 405 nm blue-violet light to crosslink and solidify the inner hydrogel solution into a non-adhesive GelMA-imid hydrogel, injecting the outer hydrogel solution and covering the outside of the solidified inner hydrogel, and irradiating it again with 405 nm blue-violet light to crosslink and solidify the outer hydrogel solution into a tissue-adhesive double-network structure hydrogel.

[0032] The beneficial effects of this invention are as follows: This invention provides an injectable bilayer drug-releasing hydrogel, its preparation method, and its application. The swelling and degradation curves and mechanical property tests of the hydrogel show that a swelling ratio of 14% for the inner hydrogel and 10% for the outer hydrogel can avoid secondary compression damage to the optic nerve. Simultaneously, it possesses suitable mechanical strength and adhesion to cope with the in vivo environment. The hydrogel precursor solution of this invention has good injectability, and the gelation time can be controlled within 2 minutes, perfectly matching the short time window of optic nerve decompression surgery, and has significant clinical practical value. It solves the major surgical risk of cerebrospinal fluid leakage during local drug administration to the optic nerve and the problems of traditional local drug delivery carriers lacking in-situ stability and being prone to displacement. It can synergistically regulate the damaged microenvironment and promote long-acting sustained release to promote nerve regeneration, successfully integrating multiple functions such as surgical safety, long-acting sustained drug release, biocompatibility, and ease of operation, providing an unprecedented and effective strategy for solving the clinical problem of optic nerve injury. Attached Figure Description

[0033] Figure 1 Transmission electron microscopy (TEM) images of MPDA and MPDA-Cle.

[0034] Figure 2 This is a schematic diagram of the crosslinking of a bilayer hydrogel system.

[0035] Figure 3 It is an injectable bilayer hydrogel system.

[0036] Figure 4 This is the SEM microstructure of the bilayer hydrogel system.

[0037] Figure 5 The swelling and degradation curves of the bilayer hydrogel system are shown.

[0038] Figure 6 This is the in vitro drug release curve of the bilayer hydrogel system.

[0039] Figure 7 For tensile property testing and burst pressure testing of bilayer hydrogel systems.

[0040] Figure 8 The graph shows the cell viability statistics and the cell live / dead staining graph.

[0041] Figure 9 This is a retinal OCT image.

[0042] Figure 10 H&E stained sections of tissue surrounding the optic nerve. Detailed Implementation

[0043] 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 a part of the embodiments of the present invention, and not all of them. 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.

[0044] Example 1: Synthesis and Characterization of Mesoporous Polydopamine Nanoparticles (MPDA)

[0045] Synthesis process:

[0046] Weigh 0.5 g of dopamine hydrochloride and 0.5 g of triblock copolymer F127 (Pluronic® F-127) and place them in a 250 mL round-bottom flask.

[0047] Add 25 mL of anhydrous ethanol and 25 mL of deionized water, place the flask on a magnetic stirrer, and stir at 500 rpm for 30 minutes at room temperature until the solid is completely dissolved.

[0048] While stirring continuously, slowly add 0.8 mL of 1,3,5-trimethylbenzene (TMB) using a pipette.

[0049] Place the flask in an ultrasonic cleaner and ultrasonically emulsify at a constant temperature of 35°C for 30 minutes until the solution turns a uniform milky white.

[0050] Transfer the flask to a 40°C constant temperature water bath and maintain stirring at 500 rpm. Slowly add 2 mL of ammonia solution (concentration: 25%) using a dropper at a rate of approximately 1 drop per second.

[0051] After the addition was complete, the flask was wrapped with aluminum foil to protect it from light, and the reaction was continued at 40°C and 500 rpm for 3 hours.

[0052] After the reaction was complete, the reaction solution was transferred to a 50 mL centrifuge tube and centrifuged for 30 minutes at 4°C and 11,000 rpm using a high-speed refrigerated centrifuge. The supernatant was then discarded.

[0053] The precipitate was washed three times with a mixture of anhydrous ethanol and deionized water (1:1, v / v), and then washed twice with a mixture of anhydrous ethanol and acetone (2:1, v / v). After each washing, the precipitate was centrifuged at 11,000 rpm for 10 minutes.

[0054] The final precipitate was freeze-dried in a freeze dryer for 24 hours to obtain black MPDA powder, which was then stored at 4°C in the dark.

[0055] Characterization results:

[0056] like Figure 1 As shown, the synthesized MPDA nanoparticles are regular spherical in shape and have obvious mesoporous structure, as observed by transmission electron microscopy (TEM, model: JEOL, Japan).

[0057] The hydrated particle size was determined to be 286.9 ± 10.5 nm and the Zeta potential was -39.1 ± 0.4 mV using a dynamic light scattering particle size analyzer (model: DLS, Malvern Instruments, UK).

[0058] Example 2: Preparation and drug loading rate determination of MPDA-loaded clomastine (MPDA@CLE)

[0059] Drug delivery process:

[0060] Accurately weigh 10 mg of the MPDA powder prepared in Example 1 and disperse it in 10 mL of chlormastine phosphate buffer solution (PBS, pH=7.4) (initial concentration of chlormastine: 1 mg / mL).

[0061] The mixture was placed in a shaker and subjected to shaking at 150 rpm for 24 hours at room temperature and in the dark.

[0062] After loading, the dispersion was transferred to a centrifuge tube and centrifuged at 4°C and 12,000 rpm for 20 minutes to collect the precipitate.

[0063] After gently washing the precipitate twice with PBS, it was freeze-dried to obtain MPDA@CLE powder.

[0064] Drug loading rate determination:

[0065] The concentration change of clomastine in the supernatant before and after loading was determined by ultraviolet-visible spectrophotometry (UV-Vis, model: UV-1780, Shimadzu, Japan), and the drug loading was calculated by standard curve.

[0066] The calculated drug loading rate of MPDA@CLE was 15.28% ± 0.13%.

[0067] Example 3: Preparation and Physicochemical Characterization of Bilayer Hydrogels

[0068] Preparation of hydrogel precursor solution:

[0069] Outer hydrogel solution (G20OD5): Weigh 200 mg of GelMA (degree of substitution DS=30) powder, dissolve it in 1 mL of PBS (pH=7.4), and shake in a 37°C water bath until completely dissolved to obtain a 20% (w / v) GelMA solution. Then add 50 mg of ODex (oxidation degree: 105.89 ± 15.88%) and 2.5 mg of LAP photoinitiator, and vortex mix thoroughly to achieve a final concentration of 20% GelMA, 5% ODex, and 0.25% LAP.

[0070] Inner layer hydrogel solution (G20imid5): Weigh 200 mg of GelMA powder and dissolve it in 1 mL of PBS, then prepare a 20% solution using the same procedure as above. Add 50 mg of 2-mercapto-1-methylimidazole and 2.5 mg of LAP, and vortex to mix thoroughly, so that the final concentration is GelMA 20%, 2-mercapto-1-methylimidazole 5%, and LAP 0.25%. If necessary, add an appropriate amount of MPDA@CLE to this solution.

[0071] Gel formation performance test:

[0072] Take 100 μL of the above solution onto a glass slide and crosslink it using a 405 nm wavelength ultraviolet light crosslinking box (light intensity: 30 mW / cm²). 2Irradiation. Observation using the inverted method showed that both formulation solutions formed stable hydrogels within 2 minutes.

[0073] Injectability and photocurability:

[0074] Two hydrogel precursor solutions can be easily injected using a 22G syringe, with a smooth and unobstructed injection process. Figure 3 As shown, under irradiation with 405 nm blue-violet light (light intensity: 10 mW / cm²), both solutions completed cross-linking within 90 seconds, forming a structurally stable hydrogel that met the time requirements for surgical procedures.

[0075] Morphological observation:

[0076] like Figure 4 As shown, the gelled hydrogel was freeze-dried, sputter-coated with gold, and observed using a scanning electron microscope (SEM, Thermo Fisher Scientific, USA). Both the inner and outer hydrogel layers exhibit a porous three-dimensional network structure.

[0077] Swelling and degradation behavior:

[0078] like Figure 5 As shown, the swelling curve was obtained by immersing the weighed dry gel (W0) in PBS and allowing it to swell to equilibrium at 37°C before weighing (W1). The swelling ratio (SR) was calculated as (W1 - W0) / W0 × 100%. The swelling ratio of the inner hydrogel was approximately 14%, and that of the outer layer was approximately 10%. This indicates that the gel is dimensionally stable in a hydrated environment and is unlikely to cause compression damage to the optic nerve tissue due to excessive swelling.

[0079] Degradation curve: In PBS containing collagenase (1 U / mL) at 37°C, a known weight of hydrogel (W) was added. i Immerse the sample in PBS containing 1 U / mL collagenase and shake at 37°C. Periodically remove the sample, blot dry, and weigh (W1). Calculate the remaining mass percentage = (W1 / W) i () × 100%. The hydrogel exhibits a slow degradation trend. The inner layer of the hydrogel completely degrades in about 300 hours, and the outer layer in about 324 hours, indicating that it can provide sustained drug release support and physical barrier function for up to several weeks in vivo.

[0080] In vitro drug release curve

[0081] The in vitro release of chlormastine was studied in PBS (pH=7.4, 37°C) using the dialysis bag method. The cumulative release rates of free chlormastine, chlormastine loaded in a monolayer hydrogel (Gel / CLE), and MPDA@CLE loaded in the bilayer hydrogel of this invention (Gel / MPDA@CLE) were compared.

[0082] like Figure 6 As shown, the bilayer hydrogel system (Gel / MPDA@CLE) of this invention exhibits a significant sustained-release effect, with the drug being released continuously for more than 168 hours (1 week) without any obvious burst release phenomenon, which meets the needs of long-acting treatment.

[0083] Biocompatibility evaluation of bilayer hydrogels

[0084] Mechanical property testing:

[0085] like Figure 7 As shown, tensile and adhesion tests were conducted using a texture analyzer (model: CTX, Brookfield Ametek, USA). The outer hydrogel G20OD5 exhibited good tensile properties and wet tissue adhesion strength. Burst pressure tests showed that it could withstand pressures exceeding 11.3 mmHg.

[0086] In vitro cytotoxicity assay (CCK-8 assay):

[0087] The rat retinal ganglion cell line (R28) was used for testing. Cells were co-cultured with standard culture medium (Control), outer layer hydrogel extract (GM-Odex), inner layer hydrogel extract (GM-imid), and inner and outer layer hydrogel drug-loaded complex (GO / GI@MPDA-Cle) for 24 h, 48 h, and 72 h, respectively.

[0088] like Figure 8 As shown in the results, there was no significant difference in cell viability between the inner and outer hydrogel extract groups compared to the Control group (p>0.05), and no obvious dead cells were observed in the inner and outer hydrogel extract groups after co-culture. This indicates that the material extract has no significant cytotoxicity.

[0089] In vivo tissue compatibility observation:

[0090] After applying the bilayer hydrogel of this invention to the optic nerve of rats, the retinal structure was observed by optical coherence tomography (OCT) at different time points, and periocular tissue was taken for histological sectioning (H&E staining).

[0091] like Figure 9 As shown, after the application of hydrogel, the structure of each layer of the retina is clear, and no obvious edema or detachment is observed.

[0092] like Figure 10 As shown, the results indicate that, compared with the sham surgery group, the hydrogel application site (Gel) showed only mild inflammatory cell infiltration, without necrosis or other organic damage, indicating that it has good in vivo tissue compatibility.

[0093] Please note to all technical personnel: Although the present invention has been described according to the specific embodiments above, the inventive concept of the present invention is not limited to this invention. Any modifications that utilize the inventive concept will be included within the scope of patent protection of this patent.

[0094] The above description is merely a preferred embodiment of the present invention. The scope of protection of the present invention is not limited to the above embodiments. All technical solutions falling within the scope of the present invention's concept are within the scope of protection of the present invention. It should be noted that for those skilled in the art, any improvements and modifications made without departing from the principles of the present invention should also be considered within the scope of protection of the present invention.

Claims

1. An injectable bilayer drug-releasing hydrogel for local drug delivery in optic nerve injury, characterized in that, The injectable bilayer drug-releasing hydrogel comprises an inner hydrogel and an outer hydrogel. The inner hydrogel is composed of methacrylamide gelatin (GelMA) and 2-mercapto-1-methylimidazole via photocrosslinking. The inner hydrogel also contains drug-encapsulated nanoparticles. The outer hydrogel is composed of methacrylamide gelatin (GelMA) and oxidized dextran (ODex) via photocrosslinking. The concentrations of methacrylamide gelatin (GelMA) and 2-mercapto-1-methylimidazole in the inner hydrogel are 20% w / v and 5% w / v, respectively. The concentrations of methacrylamide gelatin (GelMA) and oxidized dextran (ODex) in the outer hydrogel are 20% w / v and 5% w / v, respectively.

2. The injectable bilayer drug-releasing hydrogel for local drug delivery in optic nerve injury according to claim 1, characterized in that, The nanoparticles encapsulating the drug are mesoporous polydopamine nanoparticles (MPDA@CLE) encapsulated with chlormastine (CLE).

3. The injectable bilayer drug-releasing hydrogel for local drug delivery in optic nerve injury according to claim 1, characterized in that, The photocrosslinking is performed by irradiating the LAP photoinitiator with blue-violet light at a wavelength of 405 nm.

4. A method for preparing an injectable bilayer drug-releasing hydrogel for local drug delivery in optic nerve injury as described in claim 2, characterized in that, Includes the following steps: S1. Synthesis of mesoporous polydopamine nanoparticles (MPDA): Dopamine hydrochloride and triblock copolymer F127 were added to anhydrous ethanol and deionized water and magnetically stirred at room temperature until the solid was completely dissolved. Under continuous stirring, 1,3,5-trimethylbenzene (TMB) was added and ultrasonically emulsified at 35°C until the solution turned a uniform milky white color. Then, the solution was transferred to a 40°C constant temperature water bath and stirred. Ammonia water was added dropwise. After the addition was complete, the solution was protected from light and the reaction continued until the reaction was finished. The supernatant was discarded by high-speed refrigeration centrifugation. The precipitate was washed with a mixed solution of anhydrous ethanol and deionized water, and then washed with a mixed solution of anhydrous ethanol and acetone. After each washing, the solution was centrifuged. The final precipitate was freeze-dried to obtain black MPDA powder. S2. Preparation of MPDA-loaded chlormastine MPDA@CLE: MPDA powder was dispersed in chlormastine phosphate buffer solution and shaken at room temperature in the dark. After the reaction, the precipitate was collected by centrifugation, the precipitate was gently washed with PBS, and freeze-dried to obtain MPDA@CLE powder. S3. Preparation of inner layer hydrogel solution: GelMA was dissolved in PBS to prepare a 20% solution. 5% 2-mercapto-1-methylimidazole and 0.25% photoinitiator LAP were added. The mixture was stirred until homogeneous and MPDA@CLE nanoparticles loaded with chlormastine were added. Photocrosslinking was performed to obtain an inner layer hydrogel solution loaded with MPDA@CLE nanoparticles. S4. Preparation of outer hydrogel solution: Dissolve GelMA in PBS to prepare a 20% solution, add 5% ODex and 0.25% photoinitiator LAP, stir evenly, and photocrosslink to obtain the outer hydrogel solution. S5. Inject the inner hydrogel solution into the target area within the optic nerve sheath and irradiate with 405 nm blue-violet light to crosslink and solidify the inner hydrogel solution into a non-adhesive inner hydrogel. Inject the outer hydrogel solution and cover the outside of the solidified inner hydrogel. Irradiate with 405 nm blue-violet light again to crosslink and solidify the outer hydrogel solution into a double-network structure hydrogel with tissue adhesion.

5. The preparation method according to claim 4, characterized in that, In step S1, the volume ratios of anhydrous ethanol to deionized water and anhydrous ethanol to acetone are 1:1 and 2:1, respectively.

6. The preparation method according to claim 4, characterized in that, In step S2, the concentration of MPDA powder dispersed in chlormastine phosphate buffer solution is 1 mg / mL.

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