A preparation method of a gelatin hydrogel nanodelivery system for treating oral ulcer

Abstract

The present application relates to the field of biomedical materials, and discloses a preparation method of a gelatin hydrogel nano delivery system for treating oral ulcer, comprising: adding gelatin into a phosphate buffer solution, heating and stirring to obtain a gelatin matrix solution; adding polydopamine modified drug-loaded polylactic acid nanoparticles into the matrix solution under 37-45 DEG C, and slowly stirring to obtain a mixed system; adding symmetric bisuccinimidyl ester crosslinking agent dissolved in a biocompatible cosolvent into the mixed system, uniformly mixing, and in-situ crosslinking of the system to form a crosslinked gelatin hydrogel. The crosslinking agent containing disulfide bonds reacts with the amino groups of gelatin, the system is gelled without initiator, has tissue adhesion, and prolongs the residence time of the wound; the nanoparticles are covalently combined to reduce the initial drug burst of water absorption; the disulfide bonds are broken and degraded in response to the local reducing microenvironment, so that the drug release process matches the ulcer inflammation state.

Description

Technical Field

[0001] This invention relates to the field of biomedical materials technology, specifically to a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers. Background Technology

[0002] Oral ulcers are a common oral mucosal disease, often accompanied by local inflammation and pain. Currently, clinical treatment typically involves topical medication, using mouthwashes, ointments, or films to deliver anti-inflammatory drugs to the ulcer, thereby relieving symptoms and promoting mucosal healing.

[0003] Due to the unique characteristics of the oral cavity environment, existing topical drug delivery formulations have certain limitations in practical applications. The continuous secretion of saliva and mechanical movements such as chewing and speaking constantly wash away the medication adhering to the mucosa, resulting in a short retention time for traditional formulations on ulcer wounds and easy drug loss. To prolong the duration of drug action, the conventional method is to use polymeric hydrogel materials as drug carriers. However, ordinary hydrogels mainly rely on physical entanglement or non-covalent bonds. After absorbing water and swelling in the moist oral environment, their structural strength and adhesiveness easily decrease, making it difficult to maintain stable wound coverage.

[0004] Furthermore, most existing hydrogel drug delivery systems rely primarily on the passive diffusion of drug molecules within the gel matrix for release. This purely physical encapsulation delivery method is prone to burst release upon initial contact with bodily fluids, failing to achieve stable release. Simultaneously, oral ulcers possess a unique inflammatory microenvironment, typically accompanied by biochemical characteristics such as elevated local concentrations of reducing substances. Traditional drug-loaded hydrogels lack structural responsiveness to the ulcer's pathological microenvironment; the drug release rate is not regulated by the degree of inflammation at the wound site, making it difficult to achieve long-lasting, on-demand release specifically targeting the lesion, thus impacting the final efficacy of local treatment. Summary of the Invention

[0005] To address the shortcomings of existing technologies, this invention provides a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers. This method solves the problems of existing topical drug delivery formulations for oral ulcers having short retention times in the moist and mechanically frictional environment of the oral cavity, and lacking effective response to the ulcer microenvironment during drug release, making it difficult to achieve long-lasting and in-situ release, thus resulting in poor therapeutic effects.

[0006] To achieve the above objectives, the present invention provides the following technical solution: a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers, comprising the following steps: Gelatin was added to a phosphate buffer solution with a pH of 7.4 and continuously magnetically stirred under water bath heating at 37 to 50°C until completely dissolved to obtain a gelatin matrix solution with a mass fraction of 5 to 20 w / v%. At a temperature of 37 to 45°C, lyophilized polydopamine-modified drug-loaded polylactic acid nanoparticles were added to a gelatin matrix solution and stirred slowly for 10 to 30 minutes to obtain a nanoparticle-gelatin mixture system. The final mass concentration of polydopamine-modified drug-loaded polylactic acid nanoparticles in the mixture system was 1 to 10 mg / mL. The symmetrical bis(succinimide) crosslinking agent was pre-dissolved in a biocompatible co-solvent and then rapidly added to the nanoparticle gelatin mixture and quickly mixed. The final mass concentration of the symmetrical bis(succinimide) crosslinking agent in the mixture was controlled to be 5 to 20 mg / mL. The system underwent in-situ crosslinking to form a crosslinked gelatin hydrogel loaded with nanoparticles.

[0007] By employing the above technical solution, and through crosslinking a gelatin matrix with a symmetrical bis(succinimide) crosslinking agent containing disulfide bonds, and introducing polydopamine-modified drug-loaded polylactic acid nanoparticles, a delivery system with excellent biocompatibility, tissue adhesion, and microenvironment-responsive degradation and release characteristics is obtained. The specific reaction mechanism is detailed below: Step 1: Physical Dispersion At temperatures ranging from 37 to 45°C, the gelatin polypeptide chains are in a random coil state, resulting in good fluidity. The added polydopamine-modified drug-loaded polylactic acid nanoparticles possess phenolic hydroxyl and secondary amino groups on their surface. These polar groups form a hydrogen bond network with amino acid residues on the gelatin chains. This interaction lowers the surface energy of the nanoparticles, enabling them to disperse uniformly in the gelatin aqueous matrix and preventing the aggregation and sedimentation of the hydrophobic polylactic acid particles.

[0008] Step 2: In-situ covalent crosslinking The addition of a symmetrical bis(succinimide) crosslinking agent to the system initiates in-situ chemical crosslinking. The reaction mechanism is a nucleophilic substitution reaction. The gelatin macromolecular backbone contains lysine residues, and its side chains have free primary amino groups. The N-hydroxysuccinimide esters at both ends of the crosslinking agent have active ester structures. In a buffer system with a pH of 7.4, the primary amino group of the gelatin side chain undergoes a nucleophilic substitution reaction with the carbonyl carbon of the crosslinking agent, removing the N-hydroxysuccinimide leaving group and forming a stable amide bond.

[0009] The specific cross-linking reaction can be expressed as: 2 Gelatin-NH2+NHS-O-CO-RSSR-CO-O-NHS→Gelatin-NH-CO-RSSR-CO-NH-Gelatin+2 NHS-OH.

[0010] Because the crosslinking agent is a symmetrical bifunctional molecule, its two ends are bonded to different gelatin molecular chains, thereby constructing a solid three-dimensional crosslinked polymer network in the liquid system. Simultaneously, the polydopamine structure on the surface of the nanoparticles also contains a small amount of free amino groups participating in the polymerization reaction. These amino groups simultaneously participate in the aforementioned nucleophilic substitution reaction, allowing the nanoparticles to be covalently bound to the three-dimensional framework of the hydrogel, thus mitigating the burst release of the drug-loaded particles during the initial swelling of the gel matrix.

[0011] Step 3: Microenvironment-responsive degradation The cross-linked and cured hydrogel framework incorporates disulfide bonds introduced by the cross-linking agent. Oral ulcer wounds exhibit local inflammation, and the microenvironment contains high concentrations of reducing substances such as glutathione or is under oxidative stress. When this hydrogel system comes into contact with the ulcer microenvironment, the disulfide bonds in the network framework undergo reduction and breakage, i.e., the RSSR structure cleaves into two independent thiol structures R-SH. This covalent framework breakage leads to structural degradation of the hydrogel matrix, increasing gel porosity and releasing polylactic acid nanoparticles previously bound to the framework into the ulcer wound. This process achieves a match between the drug release rate and the degree of ulcer inflammation. The entire cross-linking and curing process does not require the introduction of small molecule free radical initiators or the use of ultraviolet light irradiation, ensuring the safety of the prepared system.

[0012] This invention provides a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers. It has the following beneficial effects: 1. This invention utilizes a symmetrical bis(succinimide) crosslinking agent containing disulfide bonds to undergo an in-situ nucleophilic substitution reaction with the amino groups in gelatin molecules to form a crosslinked hydrogel under conditions without ultraviolet light irradiation and additional initiators. The polar groups in the gelatin and polydopamine structures provide the hydrogel with tissue adhesion ability, enabling it to adhere to oral ulcer wounds and prolong the local retention time in a moist environment.

[0013] 2. This invention uses polydopamine to modify the surface of drug-loaded polylactic acid nanoparticles, introducing free amino groups that can participate in cross-linking reactions. During the gelation process, the nanoparticles are covalently bonded to the three-dimensional framework of the hydrogel, changing the simple physical embedding method, reducing the burst release of drugs in the early stage of gel swelling after water absorption, and helping to achieve sustained drug release.

[0014] 3. This invention introduces disulfide bonds into the hydrogel network through a crosslinking agent, giving it microenvironment-responsive degradation characteristics. When the delivery system comes into contact with the reducing inflammatory microenvironment of the oral ulcer, the disulfide bonds break, causing the hydrogel matrix to degrade and release the drug-loaded nanoparticles inside, so that the drug release process matches the pathological state of the ulcer inflammation. Attached Figure Description

[0015] Figure 1 The diagram shows the preparation of the crosslinked gelatin hydrogel of the present invention, (a) a synthesis procedure diagram of the crosslinking agent, (b) a schematic diagram of the formation of the crosslinked gelatin hydrogel, (c) a diagram showing the effect of crosslinking agent concentration on gelatin solution gelation, (d) a schematic diagram showing the mechanism of gelatin hydrogel adhesion to skin, and (e) a diagram showing the adhesion and water resistance of gelatin hydrogel on skin under artificial saliva conditions. Figure 2 The above are in vitro cumulative release kinetic curves of dexamethasone in Examples 1-3 and Comparative Example 1 of this invention; Figure 3 This is a bar chart comparing the average fluorescence intensity of RAW 264.7 macrophages and nanoparticles with different surface modifications after co-incubation for different times according to the present invention. Figure 4 This is a bar chart showing the quantitative comparison of the relative cell viability of L929 fibroblasts in vitro after culturing in hydrogel extracts of different cross-linking systems for 24 hours and 48 hours. Figure 5 Line graphs showing the quantitative statistical analysis of the relative oral ulcer area at different time points after in situ treatment in rats of each experimental group in this invention. Detailed Implementation

[0016] The technical solutions in 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 embodiments of the present invention, and not all embodiments. 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.

[0017] Preparation Examples 1-6: Preparation Example 1: This preparation example provides a method for preparing a symmetrical bis(succinimide) crosslinking agent (NHS-SS-NHS), comprising the following steps: 1 mmol of 2-hydroxyethyl disulfide was dissolved in 20 mL of anhydrous tetrahydrofuran. 2 mmol of triethylamine was added as an acid-binding agent in an ice-water bath (0 °C). An anhydrous tetrahydrofuran solution containing 0.67 mmol of triphosgene (providing a 2 mmol phosgene equivalent) was slowly added dropwise with continuous stirring. After the addition was complete, the reaction system was stirred continuously at room temperature (20 °C) for 2 hours to generate a disulfide-bonded dichloroformate intermediate. The reaction system was cooled back to 0°C, and 2 mmol of N-hydroxysuccinimide (NHS) and 2 mmol of triethylamine were added to the reaction solution containing the intermediate. The reaction was continued at room temperature (20°C) with stirring in the dark for 12 hours. After the reaction was completed, the triethylamine hydrochloride precipitate was removed by filtration, and the filtrate was concentrated by rotary evaporation. The crude product was recrystallized twice with ethyl acetate and dried under vacuum for 12 hours to obtain the purified crosslinking agent.

[0018] Preparation Example 2: This preparation example provides a method for preparing a symmetrical bis(succinimide) crosslinking agent (NHS-SS-NHS), comprising the following steps: 3 mmol of 2-hydroxyethyl disulfide was dissolved in 35 mL of anhydrous tetrahydrofuran. 6 mmol of triethylamine was added as an acid-binding agent in an ice-water bath (3 °C). An anhydrous tetrahydrofuran solution containing 2.0 mmol of triphosgene (providing a 6 mmol phosgene equivalent) was slowly added dropwise with continuous stirring. After the addition was complete, the reaction system was stirred continuously at room temperature (23 °C) for 4 hours to generate a disulfide-containing dichloroformate intermediate. The reaction system was cooled again to 3°C, and 6 mmol of N-hydroxysuccinimide (NHS) and 6 mmol of triethylamine were added to the reaction solution containing the intermediate. The reaction was continued at room temperature (23°C) with stirring in the dark for 18 hours. After the reaction was completed, the triethylamine hydrochloride precipitate was removed by filtration, and the filtrate was concentrated by rotary evaporation. The crude product was recrystallized three times with anhydrous ethanol and dried under vacuum for 18 hours to obtain the purified crosslinking agent.

[0019] Preparation Example 3: This preparation example provides a method for preparing a symmetrical bis(succinimide) crosslinking agent (NHS-SS-NHS), comprising the following steps: 5 mmol of 2-hydroxyethyl disulfide was dissolved in 50 mL of anhydrous dichloromethane. 10 mmol of triethylamine was added as an acid-binding agent in an ice-water bath (5 °C). An anhydrous dichloromethane solution containing 3.35 mmol of triphosgene (providing approximately 10 mmol of phosgene equivalent) was slowly added dropwise with continuous stirring. After the addition was complete, the reaction system was stirred continuously at room temperature (25 °C) for 6 hours to generate a disulfide-containing dichloroformate intermediate. The reaction system was cooled again to 5°C, and 10 mmol of N-hydroxysuccinimide (NHS) and 10 mmol of triethylamine were added to the reaction solution containing the intermediate. The reaction was continued at room temperature (25°C) with stirring in the dark for 24 hours. After the reaction was completed, the precipitate was removed by filtration, and the filtrate was concentrated and recrystallized three times with ethyl acetate. The filtrate was then dried under vacuum for 24 hours to obtain the purified crosslinking agent.

[0020] Preparation Example 4: This preparation example provides a method for preparing PDA-modified drug-loaded polylactic acid nanoparticles (PDA-Dex-PANs), including the following steps: 10 mg of polylactic acid (PLA) and 1 mg of dexamethasone (Dex) were dissolved in 1 mL of dichloromethane to form an organic phase. This organic phase was then slowly added dropwise to 10 mL of an aqueous solution containing 0.5 w / v% polyvinyl alcohol (PVA) under ultrasonic conditions, and emulsified for 5 minutes to form an oil-in-water emulsion. The emulsion was then continuously magnetically stirred at room temperature for 4 hours to allow complete evaporation of the organic solvent. The suspension was centrifuged and washed twice with deionized water to obtain Dex-PANs. The obtained Dex-PANs were dispersed in Tris-HCl buffer at pH 8.0 to achieve a nanoparticle concentration of 1 mg / mL. Dopamine hydrochloride monomer was added to bring the final concentration to 0.5 mg / mL, and the mixture was magnetically stirred at room temperature in the dark for 2 hours. After the reaction was completed, the nanoparticles were centrifuged (10,000 rpm, 10 minutes) and washed twice with deionized water, and then lyophilized to obtain PDA-Dex-PANs.

[0021] Preparation Example 5: This preparation example provides a method for preparing PDA-modified drug-loaded polylactic acid nanoparticles (PDA-Dex-PANs), including the following steps: 30 mg of polylactic acid (PLA) and 5 mg of dexamethasone (Dex) were dissolved in 3 mL of dichloromethane to form an organic phase. This organic phase was then slowly added dropwise to 30 mL of an aqueous solution containing 1.2 w / v% polyvinyl alcohol (PVA) under ultrasonic conditions, and emulsified for 12 minutes to form an oil-in-water emulsion. The emulsion was continuously magnetically stirred at room temperature for 8 hours to allow complete evaporation of the organic solvent. The suspension was centrifuged and washed three times with deionized water to obtain Dex-PANs. The obtained Dex-PANs were dispersed in Tris-HCl buffer at pH 8.5 to achieve a nanoparticle concentration of 5 mg / mL. Dopamine hydrochloride monomer was added to bring the final concentration to 2.5 mg / mL, and the mixture was magnetically stirred at room temperature in the dark for 7 hours. After the reaction was completed, the nanoparticles were centrifuged (12000 rpm, 15 minutes) and washed three times with deionized water, and then lyophilized to obtain PDA-Dex-PANs.

[0022] Preparation Example 6: This preparation example provides a method for preparing PDA-modified drug-loaded polylactic acid nanoparticles (PDA-Dex-PANs), including the following steps: 50 mg of polylactic acid (PLA) and 10 mg of dexamethasone (Dex) were dissolved in 5 mL of dichloromethane to form an organic phase. This organic phase was then slowly added dropwise to 50 mL of an aqueous solution containing 2.0 w / v% polyvinyl alcohol (PVA) under ultrasonic conditions, and emulsified for 20 minutes to form an oil-in-water emulsion. The emulsion was continuously magnetically stirred at room temperature for 12 hours to allow complete evaporation of the organic solvent. The suspension was centrifuged and washed three times with deionized water to obtain Dex-PANs. The obtained Dex-PANs were dispersed in Tris-HCl buffer at pH 9.0 to achieve a nanoparticle concentration of 10 mg / mL. Dopamine hydrochloride monomer was added to bring the final concentration to 5 mg / mL, and the mixture was magnetically stirred at room temperature in the dark for 12 hours. After the reaction was completed, the nanoparticles were centrifuged (15000 rpm, 20 min) and washed three times with deionized water, and then lyophilized to obtain PDA-Dex-PANs.

[0023] Examples 1-3:

[0024] Example 1: This example provides a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers, comprising the following steps: (1) Add gelatin to phosphate buffered saline (PBS) with a pH of 7.4 and stir magnetically in a 37°C water bath until completely dissolved to prepare a gelatin matrix solution with a mass fraction of 5 w / v. (2) At 37°C, the lyophilized PDA-Dex-PANs obtained in Preparation Example 4 were added to the gelatin solution above. The final mass concentration of nanoparticles in the system was controlled to be 1 mg / mL. The mixture was stirred slowly for 10 minutes to disperse it evenly, and a mixture of nanoparticles and gelatin was obtained. (3) Immediately before simulated clinical administration, take the symmetrical bissuccinimide ester crosslinking agent prepared in Preparation Example 1, dissolve it in a small amount of biocompatible cosolvent (such as a small amount of polyethylene glycol 400 or DMSO), and then quickly add it to the mixed system obtained in step (2) to control the final concentration of the crosslinking agent in the total system to 5 mg / mL. (4) The system is quickly mixed and applied to the target oral mucosa or ulcer. The NHS group of the crosslinking agent is coupled with the free amino group on the gelatin chain. The system undergoes in-situ crosslinking within a few minutes to form a crosslinked gelatin hydrogel loaded with nanoparticles.

[0025] Example 2: This example provides a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers, including the following steps: (1) Add gelatin to phosphate buffer (PBS) with a pH of 7.4 and stir magnetically in a 40°C water bath until completely dissolved to prepare a gelatin matrix solution with a mass fraction of 12.5 w / v. (2) At 40°C (to maintain good fluidity of the system), the lyophilized PDA-Dex-PANs obtained in Preparation Example 5 were added to the gelatin solution above. The final mass concentration of nanoparticles in the system was controlled to be 5 mg / mL. The mixture was stirred slowly for 20 minutes to disperse it evenly and obtain a mixture of nanoparticles and gelatin. (3) Immediately before administration, take the symmetrical bissuccinimide ester crosslinking agent prepared in Preparation Example 2, dissolve it in a trace amount of biocompatible cosolvent, and then quickly add it to the mixed system obtained in step (2), controlling the final concentration of the crosslinking agent in the total system to be 10 mg / mL. (4) The system is quickly mixed and applied to the target oral wound. The cross-linking agent in the system reacts covalently with the gelatin and the amino groups on the surface of the wound tissue. In situ gelation occurs rapidly within tens of seconds to 1 minute, forming a cross-linked gelatin hydrogel that adheres firmly to the tissue.

[0026] Example 3: This example provides a method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers, comprising the following steps: (1) Add gelatin to phosphate buffer (PBS) with a pH of 7.4 and stir magnetically in a 50°C water bath until completely dissolved to prepare a gelatin matrix solution with a mass fraction of 20 w / v. (2) At 45°C (to prevent high-concentration gelatin from prematurely undergoing physical gelation and clumping), the lyophilized PDA-Dex-PANs obtained in Preparation Example 6 were added to the above gelatin solution. The final mass concentration of nanoparticles in the system was controlled to be 10 mg / mL. The mixture was stirred slowly for 30 minutes to ensure that it was fully and uniformly dispersed, thus obtaining a mixture of nanoparticles and gelatin. (3) Immediately before administration, take the symmetrical bissuccinimide ester crosslinking agent prepared in Preparation Example 3, dissolve it in a trace amount of biocompatible cosolvent, and then quickly add it to the mixed system obtained in step (2), controlling the final concentration of the crosslinking agent in the total system to be 20 mg / mL. (4) The system is quickly mixed and applied to the target oral wound. The system undergoes high chemical cross-linking and gelation in situ within seconds, forming a high-strength, high-drug-load cross-linked gelatin hydrogel.

[0027] Comparative Examples 1-5: Comparative Example 1: Compared with Example 2, the difference is that PDA-modified dexamethasone polylactic acid nanoparticles were not used. Instead, free dexamethasone of equal drug content was directly dissolved and mixed in a gelatin matrix solution. All other aspects were the same.

[0028] Comparative Example 2: Compared with Example 2, the difference is that in step (2), bare drug-loaded polylactic acid nanoparticles (Dex-PANs) without dopamine surface modification are added, while the rest are the same.

[0029] Comparative Example 3: Compared with Example 2, the difference is that no symmetrical bissuccinimide crosslinking agent is added in step (3), and the non-covalently bonded hydrogel is formed only by physical crosslinking during the cooling process of gelatin. All other aspects are the same.

[0030] Comparative Example 4: Compared with Example 2, the difference is that the symmetrical bissuccinimide ester crosslinking agent of the present invention is not added in step (3), but is replaced by the traditional polymeric chemical crosslinking agent glutaraldehyde in equal proportion, while the rest are the same.

[0031] Comparative Example 5: Compared with Example 2, the difference is that gelatin matrix and cross-linking agent are not used. Equal amounts of PDA-Dex-PANs are directly dispersed in phosphate buffer (PBS) to prepare a simple nanoparticle suspension (simulating conventional mouthwash). All other aspects are the same.

[0032] Test Example 1: Effect of Crosslinking Agent Concentration on Gelatin Hydrogel Formation Kinetics and Structural Stability Take 6 clean standard vials, add 2 mL of the pre-prepared 5 w / v% gelatin matrix solution to each vial, and place the vials in a 37°C water bath to maintain the liquid flowability of the gelatin solution.

[0033] Using a micropipette, a co-solvent solution of the cross-linking agent (NHS-SS-NHS) was added dropwise to each of the above vials, so that the final mass concentrations of the cross-linking agent in each vial were 0 mg / mL, 1.3 mg / mL, 2.4 mg / mL, 5.2 mg / mL, 10.5 mg / mL, and 19.8 mg / mL, respectively (for ease of illustration, see attached figure). Figure 1 The label abbreviates it as nominal concentration 0, 1, 2, 5, 10, 20 mg / mL). After adding the sample, immediately vortex for 15 seconds to mix the system.

[0034] After mixing, immediately transfer the vial to a room temperature (25℃) laboratory bench and keep it upright. The gelation time was determined using the inverted vial method. Using a stopwatch, the vial was slowly inverted 180 degrees every 30 seconds initially, and every 5 seconds near the gelation point. The time point when the mixture completely stopped flowing and no droplets slid down the surface was recorded as the room temperature gelation time.

[0035] After completing the gelation time test, all vials were sealed and transferred to a 37°C incubator. The macroscopic physical phase of the samples inside the vials was observed on day 1 and day 7, respectively.

[0036] To quantify structural stability, samples were taken on day 1 and day 7, and wet weight was measured after removing surface free water. The samples were then pre-frozen at -80℃ and freeze-dried for 48 hours before being weighed dry. The remaining mass percentage (mass retention rate) of the hydrogel at each time point was calculated by comparing it with the initial theoretical polymer dry weight.

[0037] Table 1. Gelatinization time and in vitro degradation data of gelatin hydrogels at different crosslinking agent concentrations According to the data in Table 1, the concentration of the crosslinking agent directly determines the sol-gel phase transition kinetics and thermodynamic stability of the gelatin system. When the crosslinking agent concentration is in the range of 0-2.4 mg / mL, the system cannot maintain a stable three-dimensional network structure at body temperature (37℃). (See attached table.) Figure 1 The neutron diagram (c) shows the observation results of the vials in both upright and inverted positions. Samples with crosslinking agent concentrations of 0, 1.3, and 2.4 mg / mL all exhibited downward liquid flow in the inverted state at 37°C. While samples without crosslinking agent or with low crosslinking agent concentrations could undergo some physical crosslinking at room temperature (dependent on hydrogen bonds and helical structure recombination between gelatin molecules), the phase transition temperature of gelatin is typically between 30-35°C. At 37°C, thermal motion caused hydrogen bond breakage, leading to the disintegration of the physical network. This manifested as a rapid decay of remaining mass to a low level or complete liquefaction. Figure 1 The macroscopic photographs of the corresponding groups in neutron diagram (c) after cultivation on day 1 and day 7 show a transparent liquid state, confirming the disintegration process of the physical network.

[0038] As attached Figure 1 As shown in neutron diagram (a), this invention synthesizes a symmetrical bis(succinimide) ester crosslinker (NHS-SS-NHS) containing disulfide bonds via a two-step method. The N-hydroxysuccinimide (NHS) reactive ester groups at both ends of this crosslinker exhibit extremely high reactivity. (See attached diagram.) Figure 1As shown in the neutron diagram (b), the three-dimensional covalent cross-linking principle of the hydrogel network reveals that during in-situ curing, the NHS groups at both ends of the cross-linking agent undergo nucleophilic substitution reactions with the free side-chain amino groups (-NH2) of lysine residues on the gelatin polypeptide chain, forming stable covalent amide bonds and thus constructing a dense polymer cross-linking network. The establishment of this covalent cross-linking network breaks the original thermal reversibility limitation of gelatin, enabling the gel matrix to resist thermal depolymerization at physiological temperatures.

[0039] As the concentration of the crosslinking agent increased from 5.2 mg / mL to 19.8 mg / mL, the density of crosslinking sites in the system increased proportionally, and the movement of molecular chain segments was more restricted. Macroscopically, this was reflected in the gelation time decreasing from 8.3 minutes to 0.9 minutes, and the degradation resistance on day 7 also improved. The test results, combined with the attached... Figure 1 The morphological characterization evidence from the neutron diagram (c) clearly defines the critical concentration at which the crosslinking agent can exert covalent crosslinking effects in the hydrogel delivery system, confirming the rationale for using 5.2 mg / mL to 19.8 mg / mL as the concentration range for the examples.

[0040] Test Example 2: In-situ covalent adhesion and resistance to artificial saliva washout of hydrogels at tissue interfaces Fresh detached pigskin tissue was obtained, and the underlying subcutaneous fat layer and connective tissue were removed using a scalpel. Surface bloodstains were cleaned with physiological saline. The tissue was cut into 2cm × 4cm rectangular samples and soaked in pH 7.4 phosphate buffer to keep the surface moist, simulating the physiological moisture environment of the oral mucosa, before testing.

[0041] According to the parameters set in each embodiment and comparative example, the prepared gelatin matrix solution was rapidly mixed with crosslinking agent solutions of different concentrations. The final concentrations of the crosslinking agent were set to 0 mg / mL (physical control group), 1.1 mg / mL, 2.3 mg / mL, 5.4 mg / mL, 10.2 mg / mL, and 19.7 mg / mL, respectively.

[0042] Before macroscopic gelation occurs in the mixture, 0.2 mL of the mixture is quickly and evenly applied to the overlapping area of ​​two pieces of porcine skin tissue, with the overlapping area controlled at 2 cm × 2 cm. A constant vertical load of 10 N is applied to the overlapping area using weights and maintained for 10 minutes to allow the hydrogel to complete in-situ cross-linking and curing at the tissue interface.

[0043] Overlap shear strength test: The two ends of the bonded pigskin tissue were fixed in the upper and lower tensile clamps of a universal testing machine. The tensile rate was set to 5 mm / min, and the test was started until complete peeling failure occurred at the bonded surface. The maximum load peak value during the test was recorded and divided by the actual bonded area to calculate the interfacial shear strength. Each group of samples was tested in parallel for 5 times, and the average value was taken.

[0044] Artificial saliva rinsing resistance test: Take another pig skin tissue sample with the above hydrogel system coated on one side of its surface and weigh its initial total mass. Completely immerse it in an Erlenmeyer flask containing 50 mL of artificial saliva.

[0045] The conical flask was placed in a constant-temperature air bath shaker at 37°C and 100 rpm to simulate the dynamic saliva flushing environment in the oral cavity. The hydrogel shedding was observed at set time points (day 1, day 3, and day 7).

[0046] Table 2. Interfacial shear strength and resistance to artificial saliva retention of hydrogels at different crosslinking agent concentrations. According to the data in Table 2, the introduction and concentration level of the crosslinking agent have a decisive impact on the interfacial adhesion properties and resistance to liquid-phase erosion of the hydrogel. The comparative group without crosslinking agent (0 mg / mL) showed a shear strength of 4.8 kPa and peeled off the tissue surface in less than one day during the artificial saliva erosion test. The test results indicate that traditional gelatin hydrogels rely on intermolecular hydrogen bonds, van der Waals forces, and the physical interpenetrating network of polymer chains on the microscopically rough surface of the tissue to provide weak interfacial bonding. In an aqueous environment at 37°C, water molecules penetrate the interface, disrupting hydrogen bonds and leading to physical adhesion failure. Figure 1 The image recordings of the physical contact group in neutron diagram (e) show that the physically bonded hydrogels absorb water, swell, and lose their structural integrity after immersion, making them unable to resist mechanical erosion.

[0047] With increasing crosslinking agent concentration, the interfacial shear strength exhibits a non-linear growth trend. When the crosslinking agent concentration reaches above 5.4 mg / mL, the shear strength jumps to the range of 58.9 kPa to 118.5 kPa, demonstrating strong tissue adhesion. Simultaneously, under dynamic rinsing with artificial saliva for up to 7 days, hydrogel samples with concentrations in the range of 5.4-19.7 mg / mL did not experience macroscopic peeling. The underlying mechanism of this performance improvement lies in the combination of... Figure 1The neutron diagram (d) shows the principle of covalent anchoring at the tissue interface. The crosslinking agent synthesized in this invention contains highly reactive symmetrical bissuccinimide (NHS) groups. During the in-situ curing process of hydrogel, in addition to the crosslinking of amino groups inside the hydrogel matrix, the free NHS groups at the edge of the polymer network can undergo nucleophilic reactions with free amino groups (such as lysine residues) on the surface of collagen-rich tissues (such as pig skin or oral mucosa), generating dense covalent amide bonds at the hydrogel-tissue interface.

[0048] Figure 1 The neutron diagram (e) shows the results of the cross-linking of gelatin with skin group still adhering to the substrate on day 7, which corroborates the mechanical test data in Table 2. The interfacial covalent anchoring adhesion mechanism blocks the interfacial penetration of water molecules, allowing the drug delivery carrier to remain retained under the washout of saliva and mechanical friction in the oral cavity. This confirms that the system can meet the physical adhesion requirements for long-acting local oral drug delivery in clinical practice within the set process parameters.

[0049] Test Example 3: In vitro drug dual sustained-release kinetic test Take 1.0 g each of the cross-linked gelatin hydrogel samples prepared in Examples 1, 2, and 3, and the sample of Comparative Example 1 (free dexamethasone / cross-linked gelatin hydrogel).

[0050] Each group of samples was placed into a dialysis bag with a molecular weight cutoff of 3500 Da, and the two ends of the dialysis bag were sealed with a special dialysis clamp to ensure that there was no sample leakage.

[0051] Place the sealed dialysis bags into centrifuge tubes containing 50 mL of release medium. The release medium is phosphate buffer at pH 7.4, with 0.5 v / v% Tween-80 added to meet the leak conditions for dexamethasone release.

[0052] The centrifuge tube containing the sample was fixed in a constant temperature air bath shaker, with the temperature set at 37℃ and the oscillation frequency at 100 rpm.

[0053] At the designated time points (1 hour, 12 hours, day 1, day 3, day 5, day 7, day 10), accurately aspirate 1.0 mL of release solution from each centrifuge tube and immediately add 1.0 mL of fresh release medium preheated to 37°C to the original centrifuge tube.

[0054] The extracted release sample was filtered through a 0.22 μm microporous membrane, and the concentration of dexamethasone in the filtrate was determined by high performance liquid chromatography at a detection wavelength of 240 nm.

[0055] Based on the initial total drug load of each group of samples, the cumulative release percentage of dexamethasone at different time points was calculated. Each group of samples was tested in triplicate, and the average value was taken.

[0056] Table 3. Cumulative in vitro release rate (%) of dexamethasone using different delivery systems Based on the data in Table 3, combined with Figure 2 The release behavior of the drug in the hydrogel matrix is ​​controlled by both the carrier structure and the network cross-linking density. Comparative Example 1 used free dexamethasone directly embedded in a gelatin network. As a lipophilic small molecule, dexamethasone, after absorbing water and swelling in the gelatin hydrogel, diffuses outward driven by the concentration gradient due to the lack of physical barrier constraint from the hydrophobic matrix. The solid line and hollow triangle mark represent the release curve of Comparative Example 1 (free dexamethasone gel), showing a burst release effect; its cumulative release rate reached 71.2% within 1 day, and was almost completely released after 3 days, which could not meet the needs of continuous anti-inflammatory treatment for oral ulcers.

[0057] Examples 1-3 involved encapsulating dexamethasone with polylactic acid nanoparticles and modifying them with a polydopamine layer, then embedding a cross-linked gelatin network to form a dual sustained-release barrier. The dashed line and hollow circle represent the release curve of Example 1; the dotted line and hollow square represent the release curve of Example 2; and the dashed line and hollow diamond represent the release curve of Example 3. The curves of Examples 1 to 3 visually demonstrate the regulatory effect of the dual sustained-release barrier composed of polylactic acid nanoparticles and the cross-linked gelatin network on the drug release rate.

[0058] The release of a drug from within the system to the external medium involves a pathway of polylactic acid matrix degradation, permeation through a polydopamine shell, and diffusion within a high-molecular-weight cross-linked gelatin network. In the example groups, the initial release rate was controlled below 25% within the first 12 hours, suppressing burst release. Comparison of the three example groups revealed that the cross-linking density of the hydrogel modulates the release kinetics. Example 1, with its lower concentrations of cross-linking agent and gelatin, had a higher network porosity, resulting in less resistance to water molecule penetration and drug diffusion, achieving a release rate of 95.8% after 10 days. Example 3, using higher concentrations of gelatin and cross-linking agent, formed a dense covalent network structure, restricting polymer chain relaxation and water molecule penetration, leading to slow drug release; the release rate was only 69.7% after 10 days.

[0059] Example 2, representing the medium concentration group, exhibited a stable release profile, reaching 73.6% at 7 days and 86.3% at 10 days. The synergistic effect of the hydrophobic core of polylactic acid and the dense covalent network of gelatin resulted in a drug release cycle that closely matched the natural healing and remodeling cycle (7-10 days) of oral mucosal ulcers in rats / humans, confirming the feasibility of this dual sustained-release system design.

[0060] Test Example 4: Quantitative Test of Targeted Uptake Efficiency of Macrophages for Nanoparticles with Different Surface Modifications Using the fluorescent probe coumarin-6 instead of dexamethasone, polydopamine-modified fluorescent probe nanoparticles and unmodified bare fluorescent probe nanoparticles were prepared according to the preparation methods of Example 2 and Comparative Example 2, respectively.

[0061] RAW 264.7 mouse mononuclear macrophages in the logarithmic growth phase were cultured at 2 × 10⁻⁶ cells per well. 5 Cells were seeded at a density of 1000 cells per well in 6-well cell culture plates, and high-glucose DMEM medium containing 10% fetal bovine serum was added. The plates were then incubated at 37°C with 5% carbon dioxide for 24 hours until the cells adhered to the plate.

[0062] Remove the original culture medium from the wells and add fresh culture medium containing equal amounts (50 μg / mL) of polydopamine-modified nanoparticles and bare nanoparticles, respectively. Place the cell culture plates back into the constant temperature incubator and incubate for 1 hour, 2 hours and 4 hours, respectively.

[0063] After reaching the set incubation time point, the culture medium containing free nanoparticles was aspirated, and the cell surface was washed three times with phosphate buffer pre-cooled at 4°C to remove free particles that were not taken up by the cells or were not specifically physically adsorbed.

[0064] Add EDTA-free trypsin solution to each well to digest and detach the cells, then add complete culture medium to stop the digestion. Collect the cell pellet by centrifugation at 1000 rpm for 5 minutes, and resuspend it in 0.5 mL phosphate buffer to prepare a single-cell suspension.

[0065] Flow cytometry was used to analyze the cell suspensions in each group. The excitation wavelength was set to 488 nm and the emission wavelength to 525 nm. Fluorescence signals of 10,000 cellular events were collected from each sample tube, and the average fluorescence intensity of the cell community was recorded and calculated. Three replicates were set up for each group, and the average value was taken.

[0066] Table 4. Flow cytometry data (mean fluorescence intensity) of macrophage uptake of nanoparticles with different surface modifications Based on the data in Table 4, combined with Figure 3Macrophage uptake efficiency of nanoparticles showed time dependence and differences in surface modification conformation. During incubation periods of 1 hour, 2 hours, and 4 hours, the increase in cell uptake of unmodified polylactic acid nanoparticles (Comparative Example 2) was gradual. The light gray columnar structures on the left represent the cell uptake at each time point in the Comparative Example 2 (naked drug-loaded polylactic acid nanoparticles) group; the average fluorescence intensity at 4 hours was 341.6.

[0067] The polydopamine-modified nanoparticles (Example 2) showed higher detection values ​​at all time points than the comparative group. The dark gray bars on the right represent the cellular uptake at corresponding time points for the Example 2 (polydopamine-modified drug-loaded polylactic acid nanoparticles) group. Its fluorescence intensity reached 318.2 after 1 hour of incubation, close to the uptake level of the comparative group after 4 hours, and reached 895.7 after 4 hours, representing an approximately 2.6-fold increase in total uptake.

[0068] Figure 3 By comparing the differences in fluorescence signals of the two types of nanoparticles entering macrophages, the study reflects the promoting effect of the polydopamine surface modification layer on the specific endocytic uptake efficiency of macrophages. The difference in uptake efficiency is due to the chemical properties of the polydopamine coating. The surface of unmodified polylactic acid (PLA) particles is hydrophobic and lacks molecular recognition sites. Macrophage uptake of unmodified PLA particles mainly relies on non-specific macropinocytosis, resulting in limited transmembrane transport efficiency.

[0069] The method involves forming a polydopamine coating on the particle surface via oxidative self-polymerization, containing a high density of catechol and secondary amine structural groups. The catechol groups possess interfacial adhesion properties, enabling them to engage in multiple non-covalent interactions with transmembrane receptor proteins and phospholipid molecules on the macrophage plasma membrane surface, including hydrogen bonding, π-π electron stacking, and cation-π interactions. This interfacial affinity reduces the interfacial energy between the nanoparticles and the cell membrane, promoting specific anchoring of the particles on the macrophage surface and activating the receptor-mediated endocytosis pathway.

[0070] Test results demonstrate that introducing polydopamine surface modification parameters can endow the delivery system with the physicochemical ability to actively target macrophages, enabling more dexamethasone to enter the macrophage cytoplasm through the nanocarrier and be released, thereby increasing the local concentration of the drug in inflammatory effector cells and demonstrating the effectiveness of structural design in blocking the inflammatory cascade response mechanism.

[0071] Test Example 5: In vitro cytotoxicity and biosafety assessment Take 1.0 g each of the cross-linked gelatin hydrogel prepared in Example 2 and the gelatin hydrogel prepared in Comparative Example 4 using glutaraldehyde as a cross-linking agent, and sterilize them by ultraviolet irradiation for 2 hours in a clean bench.

[0072] The sterilized hydrogels were immersed in 10 mL of DMEM complete medium containing 10% fetal bovine serum and incubated at 37°C for 24 hours. The extract was then collected and filtered through a 0.22 μm sterile microporous membrane to obtain a 100% stock solution. The stock solution was diluted with fresh DMEM complete medium to prepare a 50% (v / v) diluted extract.

[0073] L929 mouse fibroblasts in the logarithmic growth phase were collected, and the cell suspension concentration was adjusted to 8 × 10⁸ cells per well. 3 Cells were seeded at a density of 100 μL in 96-well cell culture plates. 100 μL of cell suspension was added to each well, and the plates were incubated at 37°C with 5% carbon dioxide for 24 hours to allow the cells to adhere completely.

[0074] Discard the old culture medium from the 96-well plate. Divide the cells into a blank control group, Example 2 group, and Comparative Example 4 group. Add 100 μL of fresh DMEM complete culture medium to the blank control group; add 100 μL of the corresponding concentrations (100% and 50%) of the extraction solution to Example 2 group and Comparative Example 4 group, respectively. Set up 5 replicates for each group.

[0075] Return the 96-well plate to the incubator and continue culturing for 24 hours and 48 hours, respectively. After the specified time, add 10 μL of LCK-8 assay reagent to each well in the dark.

[0076] The 96-well plates were incubated for another 2 hours. The absorbance (optical density, OD value) of each well was then measured at 450 nm using a microplate reader. The average OD value of the blank control group was used as the baseline for 100% cell viability, and the relative cell viability of each experimental group was calculated.

[0077] Table 5. In vitro survival rate of L929 cells in hydrogel extracts with different cross-linking agents Based on the data in Table 5, combined with Figure 4 The chemical structure of the cross-linking agent determines the biocompatibility of the hydrogel material. The horizontal axis corresponds to five experimental groups: the blank control group (pure culture medium), the 50% and 100% concentration extract groups of Example 2, and the 50% and 100% concentration extract groups of Comparative Example 4 (conventional glutaraldehyde cross-linking). The vertical axis represents the relative cell survival rate percentage calculated based on the blank control group. Light gray bars represent cell survival rate data after 24 hours of culture; dark gray bars represent cell survival rate data after 48 hours of culture.

[0078] The comparative example 4 hydrogel, prepared using conventional glutaraldehyde crosslinking, exhibited cytotoxicity. After culturing in 100% concentration extract for 24 hours, the survival rate of L929 cells decreased to 45.2%, and further declined to 28.6% after 48 hours; even at a 50% dilution, the cell survival rate after 48 hours was only 51.4%, far below the cytotoxicity standard limit (>70%) for medical biomaterials specified by the International Organization for Standardization (ISO 10993-5).

[0079] Cytotoxicity stems from the reversibility of glutaraldehyde cross-linking and the high reactivity of unreacted aldehyde groups. When residual or free glutaraldehyde molecules enter the cell culture medium, their aldehyde groups undergo non-specific aldehyde-amine condensation reactions with proteins on the cell membrane surface and enzymes within the cell, disrupting the normal conformation and function of proteins, causing metabolic arrest and even apoptosis.

[0080] Example 2 used a synthesized symmetrical bis(succinimide) crosslinking agent containing disulfide bonds. In 50% and 100% concentration extracts, the survival rate of L929 cells remained above 90% for both 24 and 48 hours, demonstrating good biocompatibility. Figure 4 The differences in in vitro biocompatibility between the novel bis(succinimide) crosslinking agent (Example 2) and the traditional glutaraldehyde crosslinking agent (Comparative Example 4) were compared, confirming that the material had no significant in vitro cytotoxicity.

[0081] The results are supported by two chemical mechanisms. The nucleophilic substitution reaction between the NHS active ester and the gelatin amino group proceeds under mild aqueous conditions and physiological pH, with N-hydroxysuccinimide as a byproduct. This byproduct is water-soluble and non-cytotoxic, and is easily cleared through body fluid diffusion after cross-linking. The disulfide bonds introduced into the cross-linked network backbone are structurally stable under physiological conditions and do not release destructive active groups.

[0082] Furthermore, the covalent cross-linked network is stable, preventing the leakage of unreacted monomers. Test results demonstrate that using a bis(succinimide) ester structure to replace the traditional aldehyde cross-linking system eliminates the cytotoxicity induced by the cross-linking agent, ensuring that the hydrogel system applied to oral ulcer wounds does not cause secondary chemical damage to surrounding healthy mucosal tissue and newly formed epithelial cells. This confirms the safety basis of this approach in clinical application translation.

[0083] Test Example 6: Evaluation of the therapeutic effect in a rat oral ulcer model Healthy male SD rats weighing 200-250g were selected and acclimatized for one week. The rats were then generalized by intraperitoneal injection of 3% sodium pentobarbital (30mg / kg).

[0084] Mechanical damage was applied to the left buccal mucosa of rats using a sterile biopsy punch with an inner diameter of 5 mm. The damaged area was then covered with filter paper containing 50% glacial acetic acid and left for 45 seconds to induce a standardized oral mucosal chemical and mechanical ulcer model.

[0085] Rats that successfully developed the model were randomly divided into four groups: model control group (no treatment), blank gel group (only drug-free cross-linked gelatin hydrogel was applied), comparative example 1 group (free dexamethasone / cross-linked gelatin hydrogel), and example 2 group (cross-linked gelatin hydrogel loaded with polydopamine-modified nanoparticles). There were 12 rats in each group.

[0086] On day 1 after modeling (referred to as day 0), 100 μL of the corresponding mixed gel precursor solution was applied in situ to the ulcer wounds of rats according to the group, and left to stand for about 5 minutes to allow it to cross-link and solidify in situ. An equal volume of physiological saline was added to the model control group.

[0087] On days 0, 2, 4, 6 and 8 after treatment, rats were briefly anesthetized by inhalation, and the ulcer sites were photographed vertically using a digital camera with a macro lens. A standard millimeter ruler was placed next to the wound during the photography.

[0088] Using ImageJ software, the outline of the ulcer (the boundary between the yellowish-white pseudomembrane and the surrounding congested red halo) was delineated, and the actual ulcer area was calculated. The ulcer area on day 0 was counted as 100%, and the relative ulcer area at each time point was calculated. Data from 6 rats in each group were averaged.

[0089] On days 4 and 8, three rats from each group were sacrificed, and tissue samples from the ulcer sites were excised, weighed, and homogenized in homogenization buffer. The supernatant was collected by centrifugation, and the concentration of the inflammatory factor TNF-α in the local tissue was detected using an ELISA kit.

[0090] Table 6. Changes in relative area of ​​oral ulcer lesions in rats under different intervention conditions (%) Based on the data in Table 6, combined with Figure 5 The healing rate of oral ulcers in rats was highly correlated with the structure of the delivery system and the drug release kinetics. The x-axis represents the number of days of intervention; the y-axis represents the percentage of relative ulcer area calculated based on the area created on day 0 (100%). Solid lines and hollow triangles represent the natural healing curve of the model control group. In the model control group, ulcers enlarged by day 2 (106.4%), indicating increased tissue edema and necrosis due to early inflammation. Natural healing was slow by day 8, with 45.1% of the ulcer area remaining.

[0091] The dashed lines and hollow squares represent the healing curves of the blank gel group; the healing rate of the blank gel group was slightly higher than that of the model control group (31.8% residual on day 8), indicating that the hydrogel matrix can form a three-dimensional network on the wound surface through covalent cross-linking, which can act as a physical barrier to block the erosion of salivary enzymes in the oral cavity, mechanical friction from food, and bacterial infection, thus providing a stable microenvironment for tissue repair.

[0092] The dotted lines and hollow diamonds represent the healing curves of Comparative Example 1 (free dexamethasone / crosslinked gelatin hydrogel). Comparative Example 1 showed rapid ulcer area reduction (70.8%) on day 2, but the healing process stalled from day 4 to day 8, ultimately leaving 28.6% of the ulcer area. In vitro release tests confirmed this: free dexamethasone experienced a burst release after the hydrogel swelled, and the high drug concentration suppressed acute inflammation in the short term. However, due to rapid drug loss, an effective drug concentration could not be maintained during the proliferative and remodeling phases of ulcer healing, resulting in a lack of sustained anti-inflammatory effect.

[0093] The ulcer area in Example 2 group shrank faster. The dotted line and hollow circle mark represent the healing curve of Example 2 group (crosslinked gelatin hydrogel loaded with polydopamine-modified nanoparticles). On the second day, its area reduction (83.1%) was slower than that of Comparative Example 1 group, reflecting the inhibitory effect of the drug being doubly encapsulated by polylactic acid nanoparticles and crosslinked gelatin network on the initial burst release.

[0094] From day 4 onwards, the healing process in Group 2 accelerated, with the relative ulcer area decreasing to 42.4%, and reaching 5.3% by day 8, indicating that the epithelium on the wound surface had essentially completed regeneration. The in vivo treatment results validated the feasibility of the core mechanism of this approach: the covalent anchoring of the hydrogel interface ensured the continuous retention of the carrier in the dynamic oral environment; the polydopamine-modified nanoparticles released by matrix degradation could actively target and aggregate macrophages at the bottom of the wound, enter the cells through receptor-mediated endocytosis, and release dexamethasone.

[0095] Intracellular targeted drug delivery avoids the diffusion and loss of free drugs in the interstitial space and achieves high concentrations of anti-inflammatory drugs within inflammatory effector cells, interrupting the secretion cascade of pro-inflammatory factors such as TNF-α. The combination of a dual sustained-release structure and an active targeting mechanism synchronizes the drug release cycle with the physiological repair cycle of ulcer tissue, macroscopically resulting in improved wound healing quality and rate. Figure 5 By comparing the ulcer area reduction trajectories under four treatment methods, the study demonstrated the accelerating effect of the combined dual sustained-release mechanism and targeted macrophage design on the healing process of oral mucosal ulcers in vivo.

Claims

1. A method for preparing a gelatin hydrogel nanodelivery system for treating oral ulcers, characterized in that, Includes the following steps: Gelatin was added to a phosphate buffer solution with a pH of 7.4 and continuously magnetically stirred under water bath heating at 37 to 50°C until completely dissolved to obtain a gelatin matrix solution with a mass fraction of 5 to 20 w / v%. At a temperature of 37 to 45°C, lyophilized polydopamine-modified drug-loaded polylactic acid nanoparticles were added to the gelatin matrix solution and stirred slowly for 10 to 30 minutes to obtain a nanoparticle gelatin mixture system. The final mass concentration of the polydopamine-modified drug-loaded polylactic acid nanoparticles in the mixture system was 1 to 10 mg / mL. The symmetrical bissuccinimide crosslinking agent is pre-dissolved in a biocompatible co-solvent, then added to the nanoparticle gelatin mixture and rapidly mixed. The final mass concentration of the symmetrical bissuccinimide crosslinking agent in the mixture is controlled to be 5 to 20 mg / mL. The system undergoes in-situ crosslinking to form a crosslinked gelatin hydrogel loaded with the nanoparticles.

2. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 1, characterized in that, The symmetrical bis(succinimide) crosslinking agent was prepared by the following process: 2-hydroxyethyl disulfide was dissolved in anhydrous tetrahydrofuran or anhydrous dichloromethane. Triethylamine was added in an ice-water bath at 0 to 5°C, and a solution containing triphosgene was slowly added dropwise while continuously stirring. After the addition was completed, the reaction was continuously stirred at room temperature at 20 to 25°C to generate a dichloroformate intermediate containing disulfide bonds. After cooling the reaction system back to 0 to 5°C, N-hydroxysuccinimide and triethylamine were added to the reaction solution containing the dichloroformate intermediate. The reaction was continued at 20 to 25°C under light-protected stirring. After the reaction was completed, the precipitate was removed by filtration, and the filtrate was concentrated, recrystallized, and vacuum dried to obtain the symmetrical dichloroformate crosslinking agent.

3. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 2, characterized in that, The molar ratio of 2-hydroxyethyl disulfide to triphosgene is 1:0.67, the molar ratio of 2-hydroxyethyl disulfide to N-hydroxysuccinimide is 1:2, the continuous stirring reaction time for generating the disulfide-containing dichloroformate intermediate is 2 to 6 hours, and the light-protected stirring reaction time after adding the N-hydroxysuccinimide and triethylamine is 12 to 24 hours.

4. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 1, characterized in that, The polydopamine-modified drug-loaded polylactic acid nanoparticles were prepared by the following process: Polylactic acid and dexamethasone were dissolved together in dichloromethane to form an organic phase. The organic phase was slowly added dropwise to an aqueous solution containing polyvinyl alcohol under ultrasonic conditions to emulsify and form an oil-in-water emulsion. The organic solvent was then continuously stirred magnetically at room temperature to completely evaporate, resulting in a drug-loaded polylactic acid nanoparticle suspension, which was then washed with deionized water. The washed drug-loaded polylactic acid nanoparticles were dispersed in Tris-HCl buffer, and dopamine hydrochloride monomer was added and magnetically stirred at room temperature in the dark. After the reaction was completed, the nanoparticles were centrifuged, washed with deionized water, and lyophilized to obtain the polydopamine-modified drug-loaded polylactic acid nanoparticles.

5. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 4, characterized in that, The mass ratio of polylactic acid to dexamethasone is 10:1 to 5:1, the concentration of polyvinyl alcohol in the aqueous solution containing polyvinyl alcohol is 0.5 to 2.0 w / v%, the emulsification time to form the oil-in-water emulsion is 5 to 20 min, and the continuous magnetic stirring time to allow the organic solvent to evaporate completely is 4 to 12 h.

6. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 4, characterized in that, The pH of the Tris-HCl buffer is 8.0 to 9.0, the concentration of the drug-loaded polylactic acid nanoparticles in the buffer is 1 to 10 mg / mL, the final concentration of the dopamine hydrochloride monomer in the buffer is 0.5 to 5 mg / mL, and the light-protected magnetic stirring time after adding the dopamine hydrochloride monomer is 2 to 12 h.

7. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 4, characterized in that, The centrifugation process parameters are: centrifugation speed controlled between 10,000 and 15,000 r / min, and centrifugation time controlled between 10 and 20 min.

8. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 1, characterized in that, The biocompatible cosolvent is selected from either polyethylene glycol 400 or DMSO.

9. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 1, characterized in that, The gelatin was added to the phosphate buffer solution under water bath heating conditions of 40°C, and the mass fraction of the gelatin matrix solution was 12.5 w / v.

10. The method for preparing the gelatin hydrogel nanodelivery system for treating oral ulcers according to claim 1, characterized in that, The polydopamine-modified drug-loaded polylactic acid nanoparticles were added to the gelatin matrix solution at a temperature of 40°C, and the final mass concentration of the polydopamine-modified drug-loaded polylactic acid nanoparticles was 5 mg / mL. The final mass concentration of the symmetrical bis(succinimide) crosslinking agent in the mixed system is 10 mg / mL.

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