Perianal drug dressing of multi-layer conductive composite fabric, preparation method and application thereof

Abstract

The present application relates to the technical field of medical dressings, in particular to a perianal drug dressing with a multi-layer flow-guiding composite fabric, a preparation method and application thereof, the perianal drug dressing has a four-layer structure: a wet tissue adhesion layer is a polydopamine modified chitosan / oxidized hyaluronic acid hydrogel, realizing wet adhesion; a drug sustained-release layer disperses mesoporous silica loaded with ciprofloxacin hydrochloride / defensin and chitosan microspheres loaded with epidermal growth factor / honeysuckle glycoside in a polycaprolactone fiber membrane, and the drug sustained-release layer is antibacterial and healing-promoting in cooperation; a directional flow-guiding layer is a gradient-wettability sandwich structure with bionic microchannels, and the directional flow-guiding layer can realize one-way and rapid drainage; and a high-absorption outer layer is a carboxymethyl cellulose / sodium polyacrylate composite sponge, which can lock liquid efficiently, keep the wound dry, and effectively solve the clinical problems in perianal infection wound treatment, and significantly improve the treatment effect and patient comfort.

Description

Technical Field

[0001] This invention relates to the field of medical dressing technology, specifically to a perianal medicated dressing made of multi-layered diversion composite fabric, its preparation method, and its application. Background Technology

[0002] Perianal infected wounds are common and difficult-to-treat wounds in clinical practice. Their unique anatomical location and physiological environment present numerous challenges to wound management. Since the anus is the main outlet for excretion, the wound surface is frequently exposed to bacteria and other contaminants in feces, resulting in a very high risk of infection. Simultaneously, the continuous movement of the perianal area and the mechanical stress during defecation make it difficult for traditional dressings to maintain stable adhesion. Patients often need to clean the wound and change dressings frequently after defecation, increasing the workload for medical staff and causing pain and inconvenience to patients. Furthermore, perianal wounds are usually accompanied by a large amount of exudate. If exudate is not properly managed, it can lead to maceration of the surrounding skin, bacterial growth, and further delay the healing process.

[0003] Currently, dressings used clinically for perianal wound treatment mainly include traditional gauze dressings and some functional dressings. While traditional gauze dressings are inexpensive, they have drawbacks such as limited absorbency, tendency to adhere to wound tissue, and the potential for secondary trauma during dressing changes. Although some functional dressings developed in recent years, such as hydrocolloid dressings and foam dressings, have improved absorbency and comfort, they still lack a systematic design tailored to the specific needs of perianal wounds, particularly in terms of wet tissue adhesion, exudate drainage, and drug delivery.

[0004] Chinese patent CN114571807A discloses a medical fabric and its preparation method. The medical fabric adopts a three-layer structure, including a skin-friendly layer, a cross-linking layer, and a protective layer. The skin-friendly layer is made of modified cotton fiber, and the hydrophilicity and antibacterial properties of the fiber are improved through chemical grafting modification. The cross-linking layer is prepared by electrospinning of cross-linked polyimide, mainly used to provide mechanical strength. The protective layer serves as an isolation and protection layer. This technology achieves basic medical functions through a simple three-layer superposition structure, but it has the following obvious limitations: First, the technology uses a physical pressing method for interlayer bonding, and there is a lack of chemical bonding between the layers, making them prone to separation in humid environments; second, although the skin-friendly layer has been modified, it relies solely on physical contact to achieve adhesion to the tissue, making it difficult to maintain long-term adhesion in high-humidity and high-mobility areas such as the perianal region; third, the technology completely lacks an exudate management system design, and can only passively absorb the large amounts of exudate generated by perianal wounds, which can easily lead to dressing saturation, exudate leakage, and wound maceration; finally, the technology does not have a drug delivery function and cannot achieve the therapeutic goals of anti-infection and promoting healing.

[0005] From the broader perspective of medical dressing technology development, although some novel dressings with special functions have been reported, such as nanofiber dressings with unidirectional water transport capabilities and hydrogel dressings loaded with antibacterial drugs, most of these technologies only focus on achieving a single function and lack a systematic design for the specific application scenario of perianal infected wounds. Successful treatment of perianal wounds requires dressings to simultaneously possess strong adhesion, efficient exudate management capabilities, continuous antibacterial function, and the ability to promote tissue repair. This necessitates the development of a multifunctional integrated composite dressing system.

[0006] In existing technologies, wet tissue adhesion typically relies on physical adhesion mechanisms, such as van der Waals forces and mechanical interlocking. However, these physical interactions are significantly weakened in humid environments. Dopamine chemistry, inspired by mussel adhesion proteins, shows great potential in the field of wet tissue adhesion. The catechol groups in dopamine molecules can achieve strong adhesion to proteins and sugars on the tissue surface through covalent and strong hydrogen bonds. However, how to effectively integrate this mechanism into medical dressings, especially how to ensure adhesion strength while also taking into account antibacterial properties and biocompatibility, remains a technical challenge that needs to be solved.

[0007] In terms of drug delivery, traditional dressings typically employ simple drug impregnation or coating methods, resulting in a burst release initially followed by insufficient release later, which fails to meet the needs of different stages of wound healing. The treatment process for perianal infected wounds can be divided into two main stages: the acute infection phase and the proliferative remodeling phase. The former requires rapid release of high-concentration antibiotics to control infection, while the latter requires sustained release of growth factors to promote tissue repair. How to construct a drug delivery system capable of achieving this time-controlled release is currently a hot research topic and a significant challenge.

[0008] Liquid transport management is another key function of medical dressings. Traditional dressings rely on capillary action within the material itself for liquid absorption, but this passive absorption method is inefficient and prone to backflow, leading to wound re-infiltration. In recent years, biomimetic research has provided new insights into the active regulation of liquid transport. Many biological surfaces in nature possess unique directional liquid transport capabilities, such as the lubricating surface of the pitcher plant's mouth and the water-collecting structure on the back of a desert beetle. These biological surfaces typically possess asymmetric micro / nano structures and gradient wettability, enabling unidirectional liquid transport along specific directions. Applying these biomimetic principles to medical dressing design, and constructing composite structures with directional flow guidance functions, holds promise for fundamentally solving the challenge of wound exudate management.

[0009] Therefore, there is an urgent need to develop a multifunctional composite dressing specifically for perianal infected wounds. This dressing should have strong wet tissue adhesion to ensure that it does not fall off after cleaning, intelligent drug delivery function to meet the treatment needs of different healing stages, efficient directional exudate diversion system to prevent wound maceration and dressing side leakage, and also have excellent biocompatibility and healing-promoting activity. Summary of the Invention

[0010] To address the shortcomings of existing technologies, the present invention aims to provide a perianal medicated dressing made of multilayer drainage composite fabric, its preparation method, and its application. This dressing, through an innovative four-layer functional gradient structure design, organically integrates multiple functions such as wet tissue chemical adhesion, sequential drug release, directional liquid transport, and superior absorption. It can effectively solve clinical challenges in the treatment of perianal infected wounds, significantly improving treatment efficacy and patient comfort.

[0011] To achieve the above objectives, the technical solution adopted by the present invention is to provide a perianal medicated dressing with a multi-layered drainage composite fabric, which includes, from bottom to top, a wet tissue adhesion layer, a drug sustained-release layer, a directional drainage layer, and a highly absorbent outer layer. The wet tissue adhesion layer is made of a composite hydrogel membrane of polydopamine-modified quaternized chitosan and oxidized hyaluronic acid. The polydopamine-modified quaternized chitosan is obtained by grafting dopamine groups onto the quaternized chitosan molecular chain. The quaternized chitosan is prepared by reacting chitosan with a deacetylation degree of 85% to 95% with 3-chloro-2-hydroxypropyltrimethylammonium chloride at a mass ratio of 1:2 to 1:2.5. The oxidized hyaluronic acid is prepared by oxidizing sodium hyaluronate with a molecular weight of 800kDa to 1200kDa with sodium periodate at a molar ratio of 1:0.3 to 1:0.5. The polydopamine-modified quaternized chitosan and oxidized hyaluronic acid are mixed at a mass ratio of 1:1.2 to 1:1.5 and then gelled in situ through Schiff base reaction to form a hydrogel membrane with a thickness of 100μm to 150μm.

[0012] The drug-release layer comprises a mesoporous silica nanoparticle drug-loading system and a chitosan microsphere drug-loading system. The mesoporous silica nanoparticle drug-loading system loads ciprofloxacin hydrochloride and human defensin HD5, with a mass ratio of ciprofloxacin hydrochloride to human defensin HD5 of 2:1 to 3:1. The mesoporous silica nanoparticles have a particle size of 80 nm to 120 nm, a pore size of 2.5 nm to 3.5 nm, and a specific surface area greater than 900 m² / g. The chitosan microsphere drug-loading system encapsulates recombinant human epidermal cells. Chitosan microspheres contain skin growth factor and honeysuckle glycosides, with honeysuckle glycosides accounting for 5% to 8% of the dry weight of chitosan microspheres. Mesoporous silica nanoparticles and chitosan microspheres are dispersed in a polycaprolactone nanofiber membrane. The amount of mesoporous silica nanoparticles added is 3% to 5% of the mass of polycaprolactone, and the amount of chitosan microspheres added is 8% to 10% of the mass of polycaprolactone. The polycaprolactone nanofiber membrane has a fiber diameter of 300 nm to 600 nm, a porosity greater than 70%, and a thickness of 80 μm to 120 μm.

[0013] The directional flow layer is a sandwich-style gradient wettability fiber membrane structure, consisting of a hydrophobic region, a gradient transition region, and a hydrophilic region from bottom to top. The hydrophobic region is made of polyvinylidene fluoride nanofiber membrane with a fiber diameter of 200 nm to 400 nm, a water contact angle of 145° to 155°, and a thickness of 40 μm to 60 μm. The gradient transition region is made of a blend of polyvinylidene fluoride and polyvinylpyrrolidone nanofiber membrane with a thickness of 60 μm to 80 μm and a water contact angle from 1... The gradient transitions from 15° to 65°. The hydrophilic region is made of polyvinyl alcohol nanofiber membrane with a fiber diameter of 150nm to 300nm, a water contact angle of less than 15°, and a thickness of 50μm to 70μm. The surface of the gradient transition region is processed with a biomimetic microchannel array. The microchannels have an asymmetric wedge-shaped cross-section structure, a width of 80μm to 120μm, a depth of 50μm to 70μm, a spacing of 200μm to 250μm, and a microchannel density of 15 channels / cm. 2 Up to 20 strips / cm 2 .

[0014] The high-absorbency outer layer is made of a composite sponge of sodium carboxymethyl cellulose and sodium polyacrylate. The degree of substitution of sodium carboxymethyl cellulose is 0.8 to 0.9, the molecular weight of sodium polyacrylate is 2 million, the mass ratio of sodium carboxymethyl cellulose to sodium polyacrylate is 3:2, the porosity of the composite sponge is greater than 92%, the average pore size is 50 μm to 150 μm, and the dry thickness is 2 mm to 3 mm.

[0015] In a preferred embodiment of the present invention, the amount of dopamine grafted into the polydopamine-modified quaternized chitosan is 10% to 15% of the mass of the quaternized chitosan; the gelation time of the hydrogel membrane at 37°C is 8 to 12 minutes; the adhesion strength of the hydrogel membrane to pigskin is 25 kPa to 35 kPa; the moisture content of the hydrogel membrane is 75% to 85%; and the oxygen permeability is greater than 800 mL / m²·24h. The optimization of these parameters ensures that the adhesive layer has sufficient adhesion strength while maintaining good air permeability and a moist microenvironment.

[0016] Furthermore, the surface of the mesoporous silica nanoparticles is modified with disulfide-bridged polyethylene glycol (PEG) chains. The thiol density on the surface of the mesoporous silica nanoparticles is 0.8 to 1.2 thiols / nm², the molecular weight of the PEG is 2000 Da to 5000 Da, and the loading rate of ciprofloxacin hydrochloride in the mesoporous silica nanoparticles is 18% to 22%. In a simulated wound solution at pH 6.0 containing 10 mmol / L glutathione, 70% to 75% of the drug is released within 24 hours. This redox-responsive modification enables intelligent drug release in response to the inflammatory microenvironment of the wound, achieving on-demand release.

[0017] Meanwhile, the chitosan microspheres, with a particle size of 5 μm to 15 μm, were prepared by cross-linking with glutaraldehyde. The encapsulation rate of recombinant human epidermal growth factor in the chitosan microspheres was greater than 85%, and the cumulative release amount in simulated wound fluid containing 1 mg / mL of lysozyme was greater than 60% within 14 days. The enzymatic degradation characteristics of the chitosan microspheres enabled the release of growth factors to occur simultaneously with the remodeling process of wound tissue.

[0018] In another preferred embodiment of the invention, the wet tissue adhesion layer further comprises silver nanoparticles formed by in-situ reduction of silver nitrate with polydopamine. The content of the silver nanoparticles is 0.1% to 0.3% of the dry weight of the hydrogel. The wet tissue adhesion layer is further cured by cross-linking catalyzed by horseradish peroxidase, with a horseradish peroxidase concentration of 0.05 U / mL to 0.1 U / mL and a hydrogen peroxide concentration of 0.01% to 0.03%. The double cross-linked network structure endows the adhesion layer with higher mechanical strength and stability, while the introduction of silver nanoparticles provides durable broad-spectrum antibacterial activity.

[0019] The dressing of this invention integrates its layers through a hot-pressing composite process. An aqueous polyurethane dispersion is coated between the layers as an adhesive layer. The polyurethane dispersion has a solid content of 5% to 8%, a coating amount of 15 g / m² to 20 g / m², an interlayer peel strength greater than 15 N / 25 mm, and an interlayer gas permeability greater than 500 mL / m²·24 h. The total thickness of the dressing is 3.5 mm to 4.5 mm, and the basis weight is 180 g / m² to 220 g / m². This integration method ensures a strong bond between the functional layers while maintaining the overall breathability of the structure.

[0020] The present invention also provides a method for preparing the above-mentioned multilayer drainage composite fabric perianal medicated dressing, the method comprising the following steps:

[0021] First, a wet tissue adhesion layer was prepared. Chitosan with a degree of deacetylation of 85% to 95% was prepared into an alkaline solution with a mass concentration of 2 g / L to 3 g / L, and the pH was adjusted to 9 to 10. 3-chloro-2-hydroxypropyltrimethylammonium chloride was added, with a mass ratio of chitosan to 3-chloro-2-hydroxypropyltrimethylammonium chloride of 1:2 to 1:2.5. The mixture was stirred and reacted at 60°C to 70°C for 6 to 8 hours. After the reaction was completed, the mixture was purified by dialysis and freeze-dried to obtain quaternized chitosan. Quaternized chitosan was dissolved in PBS buffer at pH 8.5 to prepare a solution with a concentration of 15 g / L to 20 g / L. Dopamine hydrochloride was added, with the amount of dopamine hydrochloride added being 10% to 15% of the mass of quaternized chitosan. The mixture was reacted at room temperature under an oxygen atmosphere for 24 to 36 hours to obtain a polydopamine-modified quaternized chitosan solution. Sodium hyaluronate with a molecular weight of 800 kDa to 1200 kDa was reacted with sodium periodate at a molar ratio of 1:0.3 to 1:0.5 under light-protected conditions for 4 to 6 hours. After dialyzing, the mixture was freeze-dried to obtain oxidized hyaluronic acid. A polydopamine-modified quaternized chitosan solution was mixed with an oxidized hyaluronic acid solution at a mass ratio of 1:1.2 to 1:1.5, and the mixture was in-situ gelled at 37°C via a Schiff base reaction for 8 to 12 minutes to obtain a hydrogel membrane with a thickness of 100 μm to 150 μm.

[0022] Subsequently, a drug-release layer was prepared using hexadecyltrimethylammonium bromide as a template agent and tetraethyl orthosilicate as a silicon source. The mixture was prepared by mixing hexadecyltrimethylammonium bromide, tetraethyl orthosilicate, ammonia, and water in a molar ratio of 1:8:140:9000 and reacting at 80°C for 3 hours. The template was then removed by calcination at 650°C to obtain mesoporous silica nanoparticles. The mesoporous silica nanoparticles were then silanized using 3-mercaptopropyltrimethoxysilane, linking polyethylene glycol segments with molecular weights ranging from 2000 Da to 5000 Da via disulfide crosslinking. Ciprofloxacin hydrochloride and human defensin HD5 were mixed at a mass ratio of 2:1 to 3:1 and then loaded onto the mesoporous silica nanoparticles. Chitosan with a degree of deacetylation greater than 85% was prepared into an acetic acid solution with a concentration of 25 g / L. This solution was emulsified with liquid paraffin containing 3% Span-80 emulsifier at a volume ratio of 1:5. A glutaraldehyde crosslinking agent with a concentration of 5 g / L was added dropwise, and the reaction was carried out at 45°C for 2 hours. After washing and drying, chitosan microspheres with a particle size of 5 μm to 15 μm were obtained. Recombinant human epidermal growth factor and honeysuckle glycosides were encapsulated in the chitosan microspheres, with the amount of honeysuckle glycosides added being 5% to 8% of the dry weight of the chitosan microspheres. Polycaprolactone with a molecular weight of 80,000 Da was dissolved in a mixed solvent of chloroform and methanol in a volume ratio of 3:1 to prepare a spinning solution with a concentration of 100 g / L to 120 g / L. Drug-loaded mesoporous silica nanoparticles and chitosan microspheres were dispersed in the spinning solution, with the amount of mesoporous silica nanoparticles added being 3% to 5% of the mass of polycaprolactone, and the amount of chitosan microspheres added being 8% to 10% of the mass of polycaprolactone. Nanofiber membranes were prepared by electrospinning at a spinning voltage of 18 kV to 22 kV, a flow rate of 0.8 mL / h to 1.2 mL / h, and a receiving distance of 15 cm, resulting in a drug-release layer with a fiber diameter of 300 nm to 600 nm, a porosity greater than 70%, and a thickness of 80 μm to 120 μm.

[0023] Next, a directional flow-guiding layer was prepared. Polyvinylidene fluoride with a molecular weight of 300,000 was dissolved in a mixed solvent of N,N-dimethylformamide and acetone in a mass ratio of 6:4 to prepare a spinning solution with a concentration of 180 g / L to 200 g / L. Hydrophobic fiber membranes were prepared by electrospinning with a spinning voltage of 20 kV to 25 kV, a flow rate of 1.0 mL / h, and a receiving distance of 18 cm. Polyvinylidene fluoride nanofiber membranes with fiber diameters of 200 nm to 400 nm, water contact angles of 145° to 155°, and thicknesses of 40 μm to 60 μm were obtained. A gradient transition zone was prepared using coaxial electrospinning technology. The core layer consisted of a 150 g / L polyvinylidene fluoride solution, and the shell layer consisted of a 120 g / L polyvinylpyrrolidone solution. A continuous wettability gradient was constructed by changing the flow rate ratio of the inner and outer layers from 1:0.5 to 1:2, resulting in blended fiber membranes with a thickness of 60 μm to 80 μm and a water contact angle gradually changing from 115° to 65°. Polyvinyl alcohol with a degree of alcoholysis of 98% and a degree of polymerization of 1700 to 2000 was prepared into an aqueous solution with a concentration of 100 g / L to 120 g / L. Hydrophilic fiber membranes were prepared using electrospinning. After spinning, the fibers were crosslinked and cured with glutaraldehyde vapor at 25°C for 12 h, resulting in polyvinyl alcohol nanofiber membranes with fiber diameters of 150 nm to 300 nm, a water contact angle of less than 15°, and a thickness of 50 μm to 70 μm. A biomimetic microchannel array was fabricated on the surface of the gradient transition region using femtosecond laser micromachining technology. The microchannels have an asymmetric wedge-shaped cross-section structure with a width of 80 μm to 120 μm, a depth of 50 μm to 70 μm, and a spacing of 200 μm to 250 μm, resulting in a microchannel density of 15 channels / cm². 2 Up to 20 strips / cm 2 .

[0024] Then, a highly absorbent outer layer was prepared by mixing sodium carboxymethyl cellulose with a degree of substitution of 0.8 to 0.9 and sodium polyacrylate with a molecular weight of 2 million in an aqueous solution at a mass ratio of 3:2 to prepare a solution with a total solid content of 80 g / L to 100 g / L. 0.5% to 0.8% of N,N'-methylenebisacrylamide by monomer mass was added as a crosslinking agent, 0.3% of ammonium persulfate was added as an initiator, and 0.2% of N,N,N',N'-tetramethylethylenediamine was added as a promoter. The solution was freeze-dried at -50°C and a vacuum degree of less than 10 Pa for 48 h to form a three-dimensional porous sponge structure with a porosity greater than 92%, an average pore size of 50 μm to 150 μm, and a dry thickness of 2 mm to 3 mm.

[0025] Finally, interlayer integration was performed. The wet tissue adhesion layer, drug sustained-release layer, directional flow layer, and high-absorption outer layer were sequentially subjected to low-temperature oxygen plasma surface treatment at a plasma power of 80W for 30s. A polyurethane aqueous dispersion with a solid content of 5% to 8% was coated between each layer as an adhesive layer at a coating amount of 15g / m² to 20g / m². Interlayer integration was carried out using a hot-pressing composite process at a composite temperature of 65℃ to 75℃, a pressure of 0.3MPa to 0.5MPa, and a time of 45s to 60s, resulting in a multilayer composite dressing with a total thickness of 3.5mm to 4.5mm and a basis weight of 180g / m² to 220g / m². After being cut and shaped, the dressing was sterilized by cobalt-60 gamma irradiation at a dose of 25kGy to 30kGy.

[0026] In an improved embodiment of the preparation method, horseradish peroxidase is added during the in-situ gelation of the wet tissue adhesion layer for enzymatic cross-linking. The concentration of horseradish peroxidase is 0.05 U / mL to 0.1 U / mL. Simultaneously, hydrogen peroxide at a concentration of 0.01% to 0.03% is added as an oxidant. The reaction is carried out at 37°C for 5 to 10 minutes to achieve a dual cross-linked network structure of Schiff base cross-linking and enzyme-catalyzed cross-linking. Silver nitrate solution is added to the polydopamine-modified quaternized chitosan solution to utilize the catechol groups of polydopamine to in-situ reduce silver ions and form silver nanoparticles. The content of the silver nanoparticles is controlled to be 0.1% to 0.3% of the dry weight of the hydrogel.

[0027] Furthermore, the high-absorption outer layer can also be secondary cured using visible light-initiated thiol-olefin click chemical crosslinking. Allyl and thiol functional groups are grafted onto the molecular chains of sodium carboxymethyl cellulose and sodium polyacrylate, with a modification degree of 3 to 5 functional groups grafted per 100 monosaccharide units. Riboflavin or eosin Y is used as the photoinitiator, and the treatment is carried out under a light intensity of 5 mW / cm². 2 Up to 10mW / cm 2 Crosslinking was completed under the condition of irradiation time of 3 to 5 minutes.

[0028] The present invention also provides the application of the above-mentioned multi-layer drainage composite fabric perianal medicated dressing in the preparation of medical dressings for treating perianal infected wounds.

[0029] The beneficial effects of this invention are as follows:

[0030] This invention, through an innovative four-layer functional gradient structure design, organically integrates multiple functions such as wet tissue chemical adhesion, time-sequential drug release, directional fluid transport, and super absorption, thereby comprehensively solving the clinical challenges in the treatment of perianal infected wounds. Compared with existing technologies, this invention has the following significant advantages and unexpected technical effects.

[0031] First, the wet tissue adhesion layer of this invention employs a composite hydrogel of polydopamine-modified quaternized chitosan and oxidized hyaluronic acid, achieving strong adhesion in a humid environment through a chemical adhesion mechanism. The catechol groups in the dopamine molecule can covalently bond with functional groups such as amino, thiol, and hydroxyl groups on the tissue surface, including Schiff base reactions and Michael addition reactions. Simultaneously, multiple hydroxyl groups of catechol can form a dense hydrogen bond network with the tissue surface. This synergistic effect of covalent and non-covalent bonds results in an adhesion strength of 25 kPa to 35 kPa, far exceeding the 5 kPa to 10 kPa of traditional physically adhesive dressings. More importantly, this chemical adhesion mechanism is not weakened in water; on the contrary, the plasticizing effect of water molecules makes the polymer chains more flexible, enhancing conformational compatibility with the tissue surface. The introduction of quaternized chitosan not only endows the material with contact-type antibacterial activity, but the positive charge of the quaternary ammonium salt can electrostatically adsorb the negatively charged lipopolysaccharides and teichoic acid on the bacterial surface, disrupting the integrity of the bacterial cell membrane and causing leakage of cell contents. Simultaneously, the hydrophobic alkyl chains of the quaternary ammonium group can insert into the bacterial lipid bilayer, causing membrane structure disorder. This dual antibacterial mechanism results in an antibacterial rate exceeding 99.5%, and also enhances adhesion properties through electrostatic interactions between positively and negatively charged tissue surfaces. The aldehyde groups of oxidized hyaluronic acid can not only undergo a Schiff base reaction with the amino groups of quaternized chitosan to form a dynamic reversible cross-linked network, endowing the material with a certain self-healing ability, but also maintain a moist microenvironment in wounds. Hyaluronic acid itself has bioactivity that promotes cell migration and angiogenesis, accelerating the wound healing process.

[0032] Secondly, the drug-release layer of this invention employs a dual-carrier system of mesoporous silica nanoparticles and chitosan microspheres, achieving time-controlled release of antibiotics and growth factors. The mesoporous silica nanoparticles possess a high specific surface area and a uniform mesoporous structure, enabling efficient drug loading. More importantly, the modification of polyethylene glycol (PEG) molecular chains bridged by disulfide bonds constructs a redox-responsive gating system. Under normal physiological conditions, PEG segments cover the mesoporous surface in an extended conformation, preventing rapid drug diffusion and release. When the dressing comes into contact with the inflammatory microenvironment of the wound, the high local concentration of glutathione and reactive oxygen species triggers the breaking of disulfide bonds, causing PEG segments to detach and allowing drug molecules to rapidly diffuse out of the mesopores. This intelligent response mechanism matches antibiotic release to the severity of infection, achieving high-concentration drug release during the acute phase of infection to rapidly control bacterial proliferation, while slowing the release rate after the infection is controlled to avoid drug waste and toxic side effects. The combination of ciprofloxacin hydrochloride and human defensin HD5 further enhances the antibacterial effect. Ciprofloxacin inhibits bacterial DNA replication by suppressing bacterial DNA gyrase and topoisomerase IV, while defensin HD5 kills bacteria by disrupting the integrity of the bacterial cell membrane. Their different mechanisms of action create a synergistic effect, and defensin, as a natural antimicrobial peptide, is less likely to induce bacterial resistance. Chitosan microspheres, cross-linked with glutaraldehyde to form a stable three-dimensional network structure, can efficiently encapsulate large-molecule protein drugs such as recombinant human epidermal growth factor, with an encapsulation rate exceeding 85%. The degradation rate of chitosan microspheres is regulated by the concentration of lysozyme in wound exudate. During wound healing, with the formation of granulation tissue and the subsidence of the inflammatory response, the lysozyme level gradually increases, accelerating the enzymatic degradation of chitosan microspheres, allowing the release of epidermal growth factor and tissue remodeling to occur simultaneously. The introduction of honeysuckle glycosides not only enhances the anti-inflammatory effect, but also inhibits the activation of the NF-κB signaling pathway, downregulates the expression of pro-inflammatory factors such as TNF-α and IL-6, and reduces tissue damage caused by the inflammatory response. Simultaneously, honeysuckle glycosides promote fibroblast proliferation and collagen synthesis, forming a synergistic healing effect with epidermal growth factor. This dual-carrier time-sequential release system precisely matches the needs of different stages of wound healing, achieving a seamless transition from infection control to tissue repair.

[0033] Furthermore, the directional flow-guiding layer of this invention employs a sandwich-style gradient wettability fiber membrane combined with a biomimetic microchannel array to achieve efficient unidirectional transport of exudate. The design of this layer is inspired by the directional liquid transport structure of the pitcher plant's mouth and the back of a desert beetle. By constructing a continuous wettability gradient from hydrophobic to hydrophilic, the Laplace pressure difference drives the liquid to transport unidirectionally along the gradient direction. The hydrophobic region formed by the polyvinylidene fluoride (PVDF) nanofiber membrane has a high water contact angle of 145° to 155°, effectively repelling reverse liquid permeation. The gradient transition region formed by the PVDF and polyvinylpyrrolidone (PVP) blended nanofiber membrane achieves continuous wettability variation by changing the ratio of the two polymers, with the water contact angle smoothly transitioning from 115° to 65°. This continuous gradient avoids liquid transport resistance that may be caused by abrupt interface changes. The hydrophilic region formed by the polyvinyl alcohol (PVA) nanofiber membrane has a superhydrophilic surface angle of less than 15°, enabling rapid absorption and transport of liquid. According to the Laplace pressure formula ΔP=γ(cosθ2-cosθ1) / r, where γ is the liquid surface tension, θ1 and θ2 are the contact angles of the front and rear interfaces, and r is the capillary radius, the pressure difference generated by the wettability gradient can continuously drive the liquid to be transported from the low wettability region to the high wettability region. The high porosity and interconnected pore structure of the nanofiber membrane provide abundant capillary channels, further enhancing the liquid transport capacity. The introduction of biomimetic microchannel arrays greatly improves the directional transport efficiency. The asymmetric wedge-shaped cross-section structure makes the capillary uplift force of the liquid in the downstream direction much greater than that in the upstream direction. The small contact angle of the wedge-shaped channel on the downstream side generates a stronger capillary force, while the large contact angle on the upstream side generates resistance, forming a one-way valve effect. The size and density of the microchannels have been carefully optimized. The width of 80 μm to 120 μm and the depth of 50 μm to 70 μm provide sufficient capillary force without causing cell debris and proteins in the exudate to clog the channels due to excessive narrowness. The spacing of 200 μm to 250 μm and the density of 15 channels / cm are also optimized. 2 Up to 20 strips / cm 2 The parameters ensure sufficient transfer area while maintaining the mechanical strength of the structure. The synergistic effect of this gradient wettability and microchannel structure enables the vertical transfer velocity of exudate to exceed 2.5 cm / s, with a unidirectional transfer ratio greater than 8:1. This effectively prevents the backflow of transferred fluid, ensures that the wound surface maintains a relatively dry microenvironment, and avoids wound maceration and bacterial growth.

[0034] Fourth, the high-absorbency outer layer of this invention employs a composite sponge of sodium carboxymethyl cellulose and sodium polyacrylate, exhibiting superior liquid absorption and retention capabilities. The carboxyl groups of sodium carboxymethyl cellulose ionize in aqueous solution, generating a negative charge that forms a high-density negatively charged network with the carboxyl groups of sodium polyacrylate. According to Donnan equilibrium theory, to maintain electroneutrality, a large number of counterions and water molecules enter the polymer network, leading to extreme swelling. Simultaneously, the cross-linking effect of N,N'-methylenebisacrylamide limits the excessive extension of the polymer chains, achieving a balance between swelling pressure and the elastic recovery force of the cross-linked network, forming a stable superabsorbent hydrogel structure. The three-dimensional porous sponge structure formed during freeze-drying further enhances the liquid absorption capacity. A porosity greater than 92% and an average pore size of 50 μm to 150 μm provide a vast liquid storage space and rapid liquid transport channels. This composite sponge exhibits an absorption rate exceeding 150 g / g in physiological saline and maintains an absorption capacity of 80 g / g to 100 g / g in simulated wound exudate containing proteins and cell debris, far surpassing the 5 g / g to 10 g / g of traditional gauze dressings. More importantly, the sponge possesses excellent liquid retention properties, retaining more than 85% of its volume under 20 kPa pressure. This means that absorbed liquid will not leak out due to external pressure or dressing movement, avoiding dressing leakage and maceration of the skin around the wound. The 3:2 mass ratio of sodium carboxymethyl cellulose to sodium polyacrylate has been systematically optimized. This ratio achieves both a high absorption rate and sufficient mechanical strength, allowing the sponge to maintain its structural integrity and prevent disintegration even after full absorption.

[0035] Fifth, this invention achieves synergistic effects among functional layers through a rational interlayer integration technique. Low-temperature oxygen plasma treatment introduces active groups such as hydroxyl, carboxyl, and amino groups onto the material surface. These groups can chemically react with isocyanate groups in the polyurethane aqueous dispersion to form strong interfacial bonds. The polyurethane aqueous dispersion, acting as a bonding layer, not only provides chemical bonding but also, due to its elastic properties, buffers stress concentration caused by differences in material properties between layers, improving the overall stability of the composite structure. Precise control of temperature, pressure, and time in the hot-pressing composite process ensures tight adhesion between layers while avoiding damage to heat-sensitive drugs such as growth factors due to excessively high temperatures. An interlayer peel strength greater than 15 N / 25 mm ensures that the dressing will not separate during use, while the interlayer gas permeability is greater than 500 mL / m². 2 • 24h ensures the overall breathability of the dressing, allowing oxygen to penetrate through each layer to reach the wound surface, which is crucial for wound healing because oxygen is an essential substance for cell metabolism and collagen synthesis.

[0036] Sixth, the preparation method of this invention has a reasonable process and controllable parameters, and has good industrial feasibility. The quaternization reaction of chitosan is carried out under alkaline conditions. 3-chloro-2-hydroxypropyltrimethylammonium chloride undergoes a nucleophilic substitution reaction with the amino groups on the chitosan molecular chain, introducing quaternary ammonium groups while maintaining the linear structure and solubility of chitosan. The auto-oxidative polymerization of dopamine in an oxygen atmosphere can graft polydopamine onto the quaternized chitosan molecular chain. The catechol groups of polydopamine endow the material with adhesive function, while the melanin-like structure of polydopamine enhances the material's antioxidant capacity. The periodate oxidation of hyaluronic acid selectively opens the C2-C3 carbon-carbon bond on the sugar ring to form an aldehyde group. The light-protected conditions prevent further oxidation of the aldehyde group. The Schiff base reaction is carried out at physiological temperature and neutral pH, and the gelation time of 8 to 12 minutes meets the needs of immediate clinical use. The Stöber method for synthesizing mesoporous silica is mature. By controlling the molar ratio of reactants and reaction conditions, the particle size, pore size, and specific surface area of ​​the mesopores can be precisely controlled. The silanization reaction of 3-mercaptopropyltrimethoxysilane introduces thiol groups on the mesoporous surface, providing reaction sites for subsequent disulfide crosslinking. The emulsion crosslinking method for chitosan microspheres is simple, and the particle size and degree of crosslinking can be controlled by adjusting the emulsification conditions and crosslinking agent concentration. Electrospinning technology can prepare fiber membranes with nanoscale fiber diameters and high porosity; the fiber morphology can be precisely controlled by adjusting parameters such as spinning solution concentration, voltage, flow rate, and receiving distance. Coaxial electrospinning technology can construct core-shell structures within single fibers, providing an effective means to achieve wettability gradients. Femtosecond laser micromachining technology can achieve submicron-level precision machining, preparing complex three-dimensional microstructures without damaging the substrate material. During freeze-drying, the growth and sublimation of ice crystals form interconnected porous structures within the material; the size and distribution of pores can be controlled by adjusting the freezing rate and vacuum level. Gamma ray irradiation sterilization leaves no chemical residues in the materials. An irradiation dose of 25kGy to 30kGy can effectively kill bacterial spores and ensure the sterility of the product.

[0037] In summary, this invention, through the innovative design of a four-layer functional gradient structure and a rational manufacturing process, has successfully developed a multifunctional composite dressing specifically for the treatment of perianal infected wounds. This dressing exhibits significantly superior performance compared to existing technologies in terms of wet tissue adhesion, sequential drug release, directional liquid transport, and superior absorption. It can comprehensively address clinical challenges in the treatment of perianal wounds and has broad application prospects and significant clinical value. Detailed Implementation

[0038] The present invention will be described in detail below with reference to specific embodiments, but the scope of protection of the present invention is not limited to the following embodiments. The raw materials and reagents used in the following embodiments are all commercially available or prepared according to conventional methods, and the methods used, unless otherwise specified, are conventional methods in the art.

[0039] The main raw materials and reagents used in the examples are as follows: Chitosan was pharmaceutical grade; 3-chloro-2-hydroxypropyltrimethylammonium chloride was purchased from Shanghai Aladdin Biochemical Technology Co., Ltd.; Dopamine hydrochloride was purchased from Beijing Solarbio Technology Co., Ltd.; Sodium hyaluronate was purchased from Shandong Freda Biotechnology Co., Ltd.; Sodium periodate was purchased from Sinopharm Chemical Reagent Co., Ltd.; Hexadecyltrimethylammonium bromide was purchased from Tianjin Kemeio Chemical Reagent Co., Ltd.; Ethyl orthosilicate was purchased from Xilong Scientific Co., Ltd.; 3-mercaptopropyltrimethoxysilane was purchased from Shanghai Maclean Biochemical Technology Co., Ltd.; Polyethylene glycol was purchased from Alfaesa Chemical Co., Ltd.; Ciprofloxacin hydrochloride was purchased from Hubei Kangbaotai Fine Chemical Co., Ltd.; Human defensin HD5 was synthesized by solid-phase method from Nanjing Peptide Biotechnology Co., Ltd. Custom-synthesized reagents; recombinant human epidermal growth factor was purchased from Beijing Yiqiao Shenzhou Technology Co., Ltd.; honeysuckle glycoside was purchased from Chengdu Ruifensi Biotechnology Co., Ltd.; polycaprolactone was purchased from Sigma-Aldrich, USA; polyvinylidene fluoride was purchased from Arkema Group; polyvinylpyrrolidone was purchased from BASF, Germany; polyvinyl alcohol was purchased from Kuraray, Japan; sodium carboxymethyl cellulose, sodium polyacrylate, N,N'-methylenebisacrylamide, ammonium persulfate, and N,N,N',N'-tetramethylethylenediamine were all purchased from Shanghai Maclean Biochemical Technology Co., Ltd.; polyurethane aqueous dispersion was purchased from Wanhua Chemical Group Co., Ltd.; horseradish peroxidase was purchased from Beijing Solarbio Technology Co., Ltd.; silver nitrate was purchased from Shanghai Maclean Biochemical Technology Co., Ltd.; all other reagents were of analytical grade.

[0040] Example 1: Preparation of a perianal medicated dressing made of multilayer drainage composite fabric

[0041] The dressing prepared in this embodiment is specifically prepared using the following steps.

[0042] Step 1: Preparation of the wet tissue adhesion layer. Weigh 10g of chitosan with a degree of deacetylation of 85% and add it to 500mL of distilled water. Adjust the pH to 9 with sodium hydroxide solution and dissolve the chitosan completely under magnetic stirring to obtain an alkaline chitosan solution with a mass concentration of 2g / L. Add 20g of 3-chloro-2-hydroxypropyltrimethylammonium chloride to this solution, with a chitosan to 3-chloro-2-hydroxypropyltrimethylammonium chloride mass ratio of 1:2. Place the reaction system in a constant temperature water bath at 60℃ and react for 6 hours under mechanical stirring. During the reaction, the quaternary ammonium groups in 3-chloro-2-hydroxypropyltrimethylammonium chloride are grafted onto the amino groups of the chitosan molecular chain through nucleophilic substitution, endowing chitosan with a permanent positive charge. After the reaction is complete, transfer the product to a dialysis bag and dialyze it in flowing distilled water for 72 hours, changing the dialysate every 12 hours to thoroughly remove unreacted small molecule reagents and reaction byproducts. After dialysis, the solution was freeze-dried for 48 hours to obtain a white, fluffy quaternized chitosan powder. 5 g of quaternized chitosan was dissolved in 250 mL of pH 8.5 PBS buffer to prepare a 20 g / L solution. 0.5 g of dopamine hydrochloride was added under magnetic stirring, the amount of dopamine hydrochloride being 10% of the mass of the quaternized chitosan. The reaction flask was sealed with a rubber stopper, and oxygen was continuously introduced through a syringe. The reaction was carried out at room temperature for 24 hours. During the reaction, dopamine underwent auto-oxidative polymerization in the presence of oxygen. The resulting polydopamine was grafted onto the quaternized chitosan molecular chain via Schiff base reaction and Michael addition reaction, yielding a dark brown polydopamine-modified quaternized chitosan solution. Separately, 2g of sodium hyaluronate with a molecular weight of 800kDa was dissolved in 200mL of distilled water, and sodium periodate was added. The molar ratio of sodium hyaluronate to sodium periodate was 1:0.3. Since the average molecular weight of sodium hyaluronate is 800kDa, and the molecular weight of its repeating monosaccharide unit is approximately 400Da, 2g of sodium hyaluronate contains approximately 5mmol of repeating unit. Therefore, 1.5mmol, or approximately 0.32g, of sodium periodate needs to be added. The reaction flask was wrapped with aluminum foil to protect it from light, and the reaction was stirred at room temperature for 4h. Periodate selectively oxidizes the vicinal diol structure on the sugar ring, opening the C2-C3 carbon-carbon bond to generate an aldehyde group. After the reaction was complete, excess ethylene glycol was added to quench the remaining periodate, and then the solution was transferred to a dialysis bag and dialyzed for 48h. The solution was then freeze-dried to obtain oxidized hyaluronic acid powder. A polydopamine-modified quaternized chitosan solution and an oxidized hyaluronic acid solution were mixed at a mass ratio of 1:1.2, with the concentration of the oxidized hyaluronic acid solution adjusted to 24 g / L to ensure accurate mass ratio. The mixed solution was poured into a pre-cleaned polytetrafluoroethylene mold with a smooth bottom and placed in a constant temperature oven at 37°C. The amino groups on the quaternized chitosan and the aldehyde groups on the oxidized hyaluronic acid underwent a Schiff base reaction to form an imine crosslinking network. The solution gradually lost its fluidity and gelled in situ, with a gelation time of 8 minutes.The mixture was cured at 37°C for 2 hours to allow the crosslinking reaction to proceed fully. The hydrogel membrane was then peeled from the mold, rinsed with PBS buffer, and placed under sterile conditions for later use. The resulting hydrogel membrane had a thickness of 100 μm and exhibited good flexibility and transparency.

[0043] Step two: Preparation of the drug sustained-release layer. First, mesoporous silica nanoparticles were synthesized. 1 g of hexadecyltrimethylammonium bromide was dissolved in 480 mL of distilled water, and 3.5 mL of concentrated ammonia was added. The mixture was stirred in a 60°C water bath for 30 min until completely dissolved. Then, 8 mL of tetraethyl orthosilicate was slowly added dropwise under vigorous stirring at a rate of 0.5 mL / min. At this point, the molar ratio of hexadecyltrimethylammonium bromide, tetraethyl orthosilicate, ammonia, and water in the reaction system was 1:8:140:9000. The reaction was continued at 80°C for 3 h. Tetraethyl orthosilicate hydrolyzed and condensed to form silica, which self-assembled into a mesoporous structure under the template of hexadecyltrimethylammonium bromide micelles. After the reaction was complete, the mixture was cooled to room temperature, and the white precipitate was collected by centrifugation. The precipitate was washed three times alternately with distilled water and ethanol. The precipitate was dried in a 60℃ oven for 12 h, then placed in a muffle furnace and heated to 650℃ at a rate of 5℃ / min, and calcined at this temperature for 6 h to completely remove the organic template agent. The resulting mesoporous silica nanoparticles were white powders with a particle size of 80 nm, a pore size of 2.5 nm, and a specific surface area of ​​950 m² / g. Surface modification of the mesoporous silica nanoparticles was performed by dispersing 5 g of mesoporous silica nanoparticles in 200 mL of anhydrous toluene and sonicating for 30 min to ensure thorough dispersion. 2 mL of 3-mercaptopropyltrimethoxysilane was added, and the mixture was refluxed under nitrogen protection for 12 h. The alkoxy groups of 3-mercaptopropyltrimethoxysilane underwent a condensation reaction with the silanol groups on the mesoporous surface, introducing thiol groups into the mesoporous surface, with a thiol group density of approximately 0.8 groups / nm². The product was collected by centrifugation, washed with toluene and ethanol, and vacuum dried to obtain thiol-modified mesoporous silica. 1 g of thiol-modified mesoporous silica was dispersed in 50 mL of dimethyl sulfoxide (DMSO), and 2 g of α,ω-dithiooctanoic acid polyethylene glycol (PEG) with a molecular weight of 2000 Da was added. The mixture was stirred at room temperature for 24 h. Under alkaline conditions, the disulfide bonds on the thiol ring open and form links with the thiol groups on the mesoporous surface through disulfide bond exchange reactions, bridging the mesoporous surface with PEG segments. After the reaction was complete, the product was collected by centrifugation, thoroughly washed with DMSO and ethanol to remove unreacted PEG, and then vacuum dried to obtain PEG-modified mesoporous silica. 0.2 g of ciprofloxacin hydrochloride and 0.1 g of human defensin HD5 were dissolved in 20 mL of PBS buffer, with a mass ratio of ciprofloxacin hydrochloride to HD5 of 2:1. 1 g of PEG-modified mesoporous silica was added to the solution, and the mixture was stirred in the dark for 24 h. Drug molecules diffuse into the mesoporous channels and adsorb onto the channel surface through electrostatic interactions. Drug-loaded nanoparticles were collected by centrifugation, quickly rinsed with a small amount of water to remove the drug adsorbed on the surface, and vacuum dried to obtain drug-loaded mesoporous silica nanoparticles with a drug loading rate of 18%.

[0044] Next, chitosan microspheres were prepared. 5g of chitosan with a degree of deacetylation of 85% was dissolved in 200mL of 2% acetic acid solution to prepare a chitosan-acetic acid solution with a concentration of 25g / L. Separately, 1000mL of liquid paraffin was heated to 45℃ with 30g of Span-80 emulsifier and stirred thoroughly until dissolved. The chitosan-acetic acid solution was slowly added to the liquid paraffin at a volume ratio of 1:5. The mixture was emulsified using a high-speed disperser at 8000rpm for 30min to form a stable oil-in-water emulsion. While continuing stirring, 50mL of a 5g / L glutaraldehyde aqueous solution was added dropwise. The amount of glutaraldehyde was 5% of the chitosan mass. The two aldehyde groups of glutaraldehyde reacted with the amino groups on the chitosan molecular chain in a Schiff base reaction to form a cross-linked microsphere structure. The temperature was maintained at 45℃, and the reaction was continued for 2h. After the reaction was complete, excess petroleum ether was added to demulsify the microspheres. The microspheres were collected by centrifugation and washed repeatedly with petroleum ether, ethanol, and water to thoroughly remove liquid paraffin and emulsifier. The microspheres were dried in a 60°C oven for 12 hours to obtain a pale yellow chitosan microsphere powder with a particle size of 5 μm. 0.5 g of recombinant human epidermal growth factor and 0.25 g of honeysuckle glycoside were dissolved in 10 mL of PBS buffer, with the amount of honeysuckle glycoside added being 5% of the subsequent dry weight of the microspheres. 5 g of dried chitosan microspheres were added to the solution and soaked at 4°C for 24 hours to allow the drug to fully penetrate and adsorb into the porous structure of the microspheres. The drug-loaded microspheres were collected by centrifugation and freeze-dried under vacuum to obtain drug-encapsulated chitosan microspheres with an encapsulation rate of 85% for recombinant human epidermal growth factor.

[0045] Then, a drug-release fiber membrane was prepared. 10g of polycaprolactone with a molecular weight of 80000 Da was weighed and added to a mixed solvent of 70mL chloroform and 30mL methanol (chloroform to methanol volume ratio 3:1). The solution was heated to 40℃ under magnetic stirring to completely dissolve the polycaprolactone, yielding a spinning solution with a concentration of 100g / L. After cooling to room temperature, 0.3g of drug-loaded mesoporous silica nanoparticles and 0.8g of drug-encapsulated chitosan microspheres were added to the spinning solution. The amount of mesoporous silica nanoparticles added was 3% of the mass of polycaprolactone, and the amount of chitosan microspheres added was 8% of the mass of polycaprolactone. The solution was sonicated for 30min to ensure uniform dispersion of the drug-loaded particles and microspheres in the spinning solution. The spinning solution was then loaded into a syringe and connected to an electrospinning device. The electrospinning parameters were set as follows: voltage 18 kV, spinning solution flow rate 0.8 mL / h, distance between syringe needle and receiving plate 15 cm, receiving plate using a roller covered with aluminum foil, and rotation speed set to 100 rpm to obtain randomly oriented fiber membranes. The entire spinning process was carried out at 25℃ and 40% relative humidity. After 4 hours of spinning, the fiber membrane was peeled off from the receiving plate and dried in a vacuum drying oven for 12 hours to remove residual solvent. The resulting drug-releasing fiber membrane was a white, soft, non-woven fabric-like material with a fiber diameter of 300 nm, a porosity of 72%, and a thickness of 80 μm.

[0046] Step 3: Preparation of the directional flow-guiding layer. First, a hydrophobic fiber membrane was prepared. 18g of polyvinylidene fluoride (PVDF) with a molecular weight of 300,000 was weighed and added to a mixed solvent of 60mL N,N-dimethylformamide and 40mL acetone (mass ratio of N,N-dimethylformamide to acetone: 6:4). The mixture was stirred at 60℃ for 4 hours to completely dissolve the PVDF, resulting in a spinning solution with a concentration of 180g / L. After cooling to room temperature, electrospinning was performed with the following parameters: voltage 20kV, flow rate 1.0mL / h, and receiving distance 18cm. After 3 hours of spinning, a hydrophobic fiber membrane was obtained with a fiber diameter of 200nm, a water contact angle of 145°, and a thickness of 40μm.

[0047] Then, a gradient transition zone fiber membrane was prepared using coaxial electrospinning technology. 15g of polyvinylidene fluoride (PVDF) was dissolved in a mixed solvent of 75mL N,N-dimethylformamide and 25mL acetone to prepare a core layer spinning solution with a concentration of 150g / L. 12g of polyvinylpyrrolidone (PVP) was dissolved in 100mL ethanol to prepare a shell layer spinning solution with a concentration of 120g / L. The two spinning solutions were respectively loaded into the inner and outer syringes of a coaxial nozzle, and the flow rate was controlled by a micro-pump. Initially, the flow rate of the inner PVDF solution was 0.6mL / h, and the flow rate of the outer PPVD solution was 0.3mL / h, with an inner-outer layer flow rate ratio of 1:0.5. During the spinning process, the flow rate ratio was gradually adjusted, increasing the outer layer flow rate by 0.1mL / h every 30min until the flow rate ratio reached 1:2. This gradient change in flow rate allows the fiber to gradually transition from hydrophobic properties rich in polyvinylidene fluoride to hydrophilic properties rich in polyvinylpyrrolidone, achieving a continuous gradient in wettability. The coaxial electrospinning voltage was 22 kV, and the receiving distance was 15 cm. After 4 hours of spinning, a fiber membrane with a gradient transition zone and a thickness of 60 μm was obtained, with the water contact angle smoothly and gradually changing from 115° to 65°.

[0048] Next, a biomimetic microchannel array was fabricated on the surface of the fiber membrane in the gradient transition region. A femtosecond laser micromachining system was used, with a laser wavelength of 800 nm, a pulse width of 100 fs, and a repetition rate of 1 kHz. The distribution pattern of the microchannels was designed using computer-aided design. The microchannels are arranged in parallel, with a width of 80 μm. An etching depth of 50 μm was achieved by controlling the laser scanning speed and energy density. The microchannel cross-section was designed as an asymmetric wedge shape, with a 30° co-current sidewall angle and a 60° counter-current sidewall angle. This asymmetric structure was achieved by changing the laser incident angle and power. The spacing between the microchannels was 200 μm, and the microchannel density was 15 channels / cm². 2 The laser processing is carried out under nitrogen protection to prevent material oxidation. After processing, the surface is ultrasonically cleaned with isopropanol to remove processing debris, and then vacuum dried to obtain a fiber membrane with a gradient transition zone containing a biomimetic microchannel array.

[0049] Finally, a hydrophilic fiber membrane was prepared. 10g of polyvinyl alcohol (PVA) with a degree of hydrolysis of 98% and a degree of polymerization of 1700 was weighed and added to 100mL of distilled water. The solution was stirred at 90℃ for 3 hours to completely dissolve, yielding a 100g / L PVA aqueous solution. After cooling to room temperature, electrospinning was performed with the following parameters: voltage 15kV, flow rate 0.5mL / h, and receiving distance 12cm. Due to the high surface tension of the aqueous solution, the spinning process needed to be carried out in an environment with a relative humidity of less than 30% to prevent fiber breakage. After 3 hours of spinning, a PVA fiber membrane was obtained. The membrane was placed in a sealed container with a petri dish containing a glutaraldehyde aqueous solution at the bottom. It was then vapor-crosslinked and cured at 25℃ for 12 hours. The glutaraldehyde vapor reacted with the hydroxyl groups on the PVA molecular chains to form acetal crosslinks, improving the water resistance of the fiber membrane. After crosslinking, the membrane was rinsed with plenty of water to remove residual glutaraldehyde and then vacuum-dried to obtain the hydrophilic fiber membrane with a fiber diameter of 150nm, a water contact angle of less than 15°, and a thickness of 50μm.

[0050] Step 4: Preparation of the high-absorbency outer layer. Weigh 6g of sodium carboxymethyl cellulose with a degree of substitution of 0.8 and 4g of sodium polyacrylate with a molecular weight of 2 million, in a mass ratio of 3:2. Add them to 100mL of distilled water and dissolve them completely under magnetic stirring to obtain a homogeneous and transparent solution with a total solids content of 100g / L. Add 50mg of N,N'-methylenebisacrylamide as a crosslinking agent, at a dosage of 0.5% of the total monomer mass. Add 30mg of ammonium persulfate as an initiator, at a dosage of 0.3% of the total monomer mass, and add 20mg of N,N,N',N'-tetramethylethylenediamine as an accelerator, at a dosage of 0.2% of the total monomer mass. Stir thoroughly to ensure uniform mixing of all components, then pour the solution into a polypropylene mold pre-cooled to -20℃. The mold dimensions are 10cm × 10cm × 1cm. The mold was quickly transferred to a freeze dryer pre-cooled to -50°C and freeze-dried for 48 hours under a vacuum of less than 10 Pa. During the freezing process, water molecules crystallized to form ice crystals. The growth of these ice crystals displaced polymer chains to the grain boundary regions, creating areas of polymer enrichment. In the subsequent sublimation drying process, the ice crystals sublimated, leaving pores and forming a three-dimensional interconnected porous structure. After freeze-drying, the sponge was removed from the mold, yielding a white, fluffy, highly absorbent outer layer material with a porosity of 92%, an average pore size of 50 μm, and a dry thickness of 2 mm.

[0051] Step 5: Layer integration. The prepared wet tissue adhesion layer, drug sustained-release layer, directional flow layer, and high-absorption outer layer are sequentially subjected to low-temperature oxygen plasma surface treatment to enhance the interfacial bonding. The plasma treatment equipment uses an RF power supply with a power setting of 80W, oxygen as the treatment gas with a flow rate of 50mL / min, a vacuum degree of 50Pa, and a treatment time of 30s for each layer. High-energy particles in the plasma bombard the material surface, breaking surface chemical bonds and introducing active groups such as hydroxyl and carboxyl groups, while simultaneously increasing surface roughness. A polyurethane aqueous dispersion with a solid content of 5% is uniformly coated onto the surface of the drug sustained-release layer using a micropipette, with the coating amount controlled at 15g / m². Then, the wet tissue adhesion layer and the drug sustained-release layer are aligned and stacked, and placed in a hot press. The hot pressing temperature is set to 65℃, the pressure to 0.3MPa, and the holding time to 45s. During the hot-pressing process, the isocyanate groups in the polyurethane molecules react with the active groups on the material surface to form chemical bonds. Simultaneously, appropriate temperature and pressure ensure a tight bond between the two layers. A directional flow-guiding layer and a high-absorption outer layer are laminated sequentially in the same manner, with a polyurethane aqueous dispersion applied to the lower layer surface before each lamination. After all laminations are completed, the multilayer dressing is cured in a 60℃ oven for 2 hours to ensure full bonding between the layers. The dressing is then cut into 10cm × 10cm squares, and each piece is individually sealed in a medical-grade aluminum foil bag. The packaged dressings are then sent to an irradiation center for sterilization using cobalt-60 gamma rays at a dose of 25 kGy. Gamma rays have extremely strong penetrating power, capable of killing bacteria and spores inside the dressing without leaving any chemical residues. After sterilization, the dressings undergo sterility testing to confirm compliance with medical device sterilization standards before being stored in the warehouse. The resulting multilayer composite dressing has a total thickness of 3.5 mm and a basis weight of 180 g / m².

[0052] Example 2: Preparation of a perianal medicated dressing made of multi-layered drainage composite fabric

[0053] The dressing prepared in this embodiment incorporates both horseradish peroxidase-catalyzed cross-linking and nano-silver antibacterial functions. The specific preparation method is basically the same as in Example 1, with the following differences.

[0054] In step one, the degree of deacetylation of chitosan was selected to be 90%, and an alkaline solution with a mass concentration of 2.5 g / L was prepared. The mass ratio of chitosan to 3-chloro-2-hydroxypropyltrimethylammonium chloride was 1:2.25, the reaction temperature was 65℃, and the reaction time was 7 h. The amount of dopamine hydrochloride added was 12.5% ​​of the mass of quaternized chitosan, and the concentration of the quaternized chitosan solution was 17.5 g / L. The reaction was carried out at room temperature under an oxygen atmosphere for 30 h. Sodium hyaluronate with a molecular weight of 1000 kDa was selected, and the molar ratio with sodium periodate was 1:0.4. The reaction time was 5 h. The polydopamine-modified quaternized chitosan solution and the oxidized hyaluronic acid solution were mixed at a mass ratio of 1:1.35. Horseradish peroxidase with a concentration of 0.075 U / mL was added to the mixed solution, and hydrogen peroxide with a concentration of 0.02% was added as an oxidant. After thorough mixing, the mixture was poured into a mold and incubated at 37°C for 10 minutes for in-situ gelation. During gelation, horseradish peroxidase catalyzed the oxidative coupling reaction of the catechol groups on polydopamine, forming an additional covalent cross-linked network that cross-links with Schiff bases to form a double cross-linked system. Furthermore, silver nitrate solution was added to the polydopamine-modified quaternized chitosan solution, with the amount of silver nitrate controlled to ensure that the final hydrogel contained 0.2% silver nanoparticles by dry weight. The catechol groups of polydopamine have reducing properties, capable of reducing silver ions to silver nanoparticles, while polydopamine acts as a stabilizer to prevent the aggregation of silver nanoparticles. The resulting hydrogel film was 125 μm thick, light gray in color, and had a smooth and delicate surface.

[0055] In step two, the synthesis reaction temperature of mesoporous silica nanoparticles was maintained at 80℃ for 3 hours, followed by calcination at 650℃ for 6 hours, yielding mesoporous silica nanoparticles with a particle size of 100 nm, a pore size of 3 nm, and a specific surface area of ​​925 m² / g. During surface modification, the amount of 3-mercaptopropyltrimethoxysilane was appropriately increased to achieve a thiol density of 1.0 thiols / nm². The molecular weight of polyethylene glycol was selected as 3500 Da. The mass ratio of ciprofloxacin hydrochloride to human defensin HD5 was adjusted to 2.5:1, with a drug loading rate of 20%. In the preparation of chitosan microspheres, the chitosan solution concentration was maintained at 25 g / L, the crosslinking reaction temperature was 45℃, and the time was 2 hours, yielding chitosan microspheres with a particle size of 10 μm. The amount of honeysuckle glycoside added was 6.5% of the dry weight of the chitosan microspheres, achieving an encapsulation rate of 87% for recombinant human epidermal growth factor. In preparing the fibrous membrane, the polycaprolactone spinning solution concentration was 110 g / L, the amount of drug-loaded mesoporous silica nanoparticles added was 4% of the mass of polycaprolactone, and the amount of chitosan microspheres added was 9%. The electrospinning parameters were: voltage 20 kV, flow rate 1.0 mL / h, and receiving distance 15 cm. A drug-release fibrous membrane with a fiber diameter of 450 nm, a porosity of 75%, and a thickness of 100 μm was obtained.

[0056] In step three, the hydrophobic fiber membrane was prepared using a polyvinylidene fluoride (PVDF) spinning solution with a concentration of 190 g / L, a spinning voltage of 22.5 kV, a flow rate of 1.0 mL / h, and a receiving distance of 18 cm, resulting in a fiber membrane with a diameter of 300 nm, a water contact angle of 150°, and a thickness of 50 μm. In the coaxial spinning of the gradient transition zone, the core layer PVDF solution concentration was 150 g / L, and the shell layer polyvinylpyrrolidone (PVP) solution concentration was 120 g / L. The flow rate ratio was gradually adjusted from 1:0.5 to 1:2, resulting in a fiber membrane with a thickness of 70 μm and a water contact angle gradually decreasing from 115° to 65°. The femtosecond laser micromachining created microchannels with a width of 100 μm, a depth of 60 μm, a spacing of 225 μm, and a microchannel density of 17.5 channels / cm. 2 The hydrophilic fiber membrane was prepared using a polyvinyl alcohol solution with a concentration of 110 g / L and a degree of polymerization of 1850. After spinning, the fiber was cross-linked with glutaraldehyde vapor for 12 h to obtain a fiber membrane with a fiber diameter of 225 nm, a water contact angle of less than 15°, and a thickness of 60 μm.

[0057] In step four, sodium carboxymethyl cellulose (CCMC) has a degree of substitution of 0.85. CCMC and sodium polyacrylate are mixed at a mass ratio of 3:2, resulting in a total solid content of 90 g / L. The amount of N,N'-methylenebisacrylamide used is 0.65% of the total monomer mass, ammonium persulfate is 0.3%, and N,N,N',N'-tetramethylethylenediamine is 0.2%. The freeze-drying temperature is -50℃, the vacuum degree is less than 10 Pa, and the time is 48 h, yielding a high-absorbency outer layer with a porosity of 95%, an average pore size of 100 μm, and a dry thickness of 2.5 mm.

[0058] In step five, the plasma treatment power was 80W for 30 seconds. The solid content of the polyurethane aqueous dispersion was 6.5%, and the coating weight was 17.5 g / m². The hot-pressing composite temperature was 70℃, the pressure was 0.4 MPa, and the time was 52.5 seconds. The irradiation dose was 27.5 kGy. The resulting multilayer composite dressing had a total thickness of 4 mm and a basis weight of 200 g / m².

[0059] Example 3: Preparation of a perianal medicated dressing made of multi-layered drainage composite fabric

[0060] The dressing prepared in this embodiment is prepared in a method that is basically the same as that in Example 1, except for the following points.

[0061] In step one, the degree of deacetylation of chitosan was selected to be 95%, and an alkaline solution with a mass concentration of 3 g / L was prepared. The pH value was adjusted to 10. The mass ratio of chitosan to 3-chloro-2-hydroxypropyltrimethylammonium chloride was 1:2.5. The reaction temperature was 70℃, and the reaction time was 8 h. The amount of dopamine hydrochloride added was 15% of the mass of quaternized chitosan, and the concentration of the quaternized chitosan solution was 15 g / L. The reaction was carried out at room temperature under an oxygen atmosphere for 36 h. The molecular weight of sodium hyaluronate was selected to be 1200 kDa, and the molar ratio of sodium periodate was 1:0.5. The reaction time was 6 h. The polydopamine-modified quaternized chitosan solution and the oxidized hyaluronic acid solution were mixed at a mass ratio of 1:1.5. The concentration of horseradish peroxidase was 0.1 U / mL, and the concentration of hydrogen peroxide was 0.03%. The reaction was carried out at 37℃ for 12 min for in-situ gelation. The amount of silver nitrate added was controlled to ensure that the final hydrogel contained 0.3% silver nanoparticles by dry weight. The resulting hydrogel film had a thickness of 150 μm.

[0062] In step two, the mesoporous silica nanoparticles had a particle size of 120 nm, a pore size of 3.5 nm, and a specific surface area of ​​980 m² / g. The thiol density was 1.2 thiols / nm², and the polyethylene glycol molecular weight was 5000 Da. The mass ratio of ciprofloxacin hydrochloride to human defensin HD5 was 3:1, with a drug loading rate of 22%. The chitosan microspheres had a particle size of 15 μm, and the amount of honeysuckle glycoside added was 8% of the dry weight of the chitosan microspheres, achieving an encapsulation rate of 90% for recombinant human epidermal growth factor. The polycaprolactone spinning solution concentration was 120 g / L, the amount of drug-loaded mesoporous silica nanoparticles added was 5% of the mass of polycaprolactone, and the amount of chitosan microspheres added was 10%. The electrospinning parameters were: voltage 22 kV, flow rate 1.2 mL / h, and receiving distance 15 cm. A drug-releasing fiber membrane with a fiber diameter of 600 nm, a porosity of 78%, and a thickness of 120 μm was obtained.

[0063] In step three, the hydrophobic fiber membrane was prepared using a 200 g / L polyvinylidene fluoride spinning solution, a spinning voltage of 25 kV, a flow rate of 1.0 mL / h, and a receiving distance of 18 cm, resulting in a fiber membrane with a diameter of 400 nm, a water contact angle of 155°, and a thickness of 60 μm. The gradient transition region had a thickness of 80 μm, with the water contact angle gradually changing from 115° to 65°. The microchannels had a width of 120 μm, a depth of 70 μm, a spacing of 250 μm, and a microchannel density of 20 channels / cm. 2 The hydrophilic fiber membrane was prepared using a polyvinyl alcohol solution with a concentration of 120 g / L and a degree of polymerization of 2000, resulting in a fiber membrane with a fiber diameter of 300 nm, a water contact angle of less than 15°, and a thickness of 70 μm.

[0064] In step four, the degree of substitution of sodium carboxymethyl cellulose is 0.9, and the total solid content is 80 g / L. The amount of N,N'-methylenebisacrylamide used is 0.8% of the total monomer mass, resulting in a high-absorption outer layer with a porosity of 98%, an average pore size of 150 μm, and a dry thickness of 3 mm.

[0065] In step five, the polyurethane aqueous dispersion has a solid content of 8% and a coating weight of 20 g / m². The hot-pressing composite temperature is 75℃, the pressure is 0.5 MPa, and the time is 60 s. The irradiation dose is 30 kGy. The resulting multilayer composite dressing has a total thickness of 4.5 mm and a basis weight of 220 g / m².

[0066] Example 4: Preparation of a perianal medicated dressing made of multilayer drainage composite fabric

[0067] This embodiment verifies the effectiveness of using visible light-initiated thiol-ene click chemical crosslinking for secondary curing of the high-absorption outer layer. The preparation method is basically the same as in Example 2, except for the preparation steps of the high-absorption outer layer.

[0068] In step four, sodium carboxymethyl cellulose and sodium polyacrylate were first functionalized. 6g of sodium carboxymethyl cellulose was dissolved in 50mL of distilled water, and 0.3g of allyl glycidyl ether and 0.1g of triethylamine catalyst were added. The mixture was reacted at 60℃ for 12h. The epoxy groups of the allyl glycidyl ether reacted with the hydroxyl groups of the sodium carboxymethyl cellulose, grafting allyl groups onto the molecular chain. The degree of modification was approximately 4 allyl groups grafted per 100 monosaccharide units. After dialysis purification, the mixture was freeze-dried to obtain allyl-modified sodium carboxymethyl cellulose. 4g of sodium polyacrylate was dissolved in 50mL of distilled water, and 0.5g of cysteine ​​hydrochloride and 0.5g of 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide hydrochloride were added as coupling agents. The mixture was reacted at room temperature for 24h. The amino group of cysteine ​​was linked to the carboxyl group of sodium polyacrylate via an amide bond, introducing thiol groups into the molecular chain. The degree of modification was approximately 4 thiol groups grafted per 100 monosaccharide units. After dialysis purification, lyophilized sodium polyacrylate was obtained. Allyl-modified sodium carboxymethyl cellulose and thiol-modified sodium polyacrylate were mixed at a mass ratio of 3:2 and dissolved in 100 mL of distilled water. 0.05 g of riboflavin was added as a photoinitiator, and the mixture was stirred thoroughly until dissolved. Then, a crosslinking agent, initiator, and accelerator were added according to the method in Example 2, and the mixture was poured into a mold for lyophilization. After lyophilization, the sponge was irradiated under a visible light LED lamp with a wavelength of 450 nm and a light intensity of 7.5 mW / cm². 2The irradiation time was 4 minutes. Under visible light irradiation, riboflavin was excited to generate free radicals, initiating a thiol-ene click reaction between thiol and allyl groups, forming an additional cross-linked network. Photo-initiated cross-linking has spatial selectivity and time controllability, and can further improve the cross-linking density and mechanical strength of the material without destroying the pre-formed cross-linked network. Other steps were the same as in Example 2, and the resulting multilayer composite dressing had a total thickness of 4 mm and a basis weight of 200 g / m².

[0069] Example 5: Preparation of a perianal medicated dressing made of multi-layered drainage composite fabric

[0070] This embodiment uses different parameter combinations to verify the stability of the preparation process and the reproducibility of product performance. The degree of deacetylation of chitosan is 88%, the mass concentration is 2.3 g / L, the mass ratio of chitosan to 3-chloro-2-hydroxypropyltrimethylammonium chloride is 1:2.3, the reaction temperature is 68℃, and the reaction time is 7.5 h. The dopamine grafting amount is 13% of the mass of quaternized chitosan, and the reaction time is 33 h. The molecular weight of sodium hyaluronate is 950 kDa, the molar ratio of sodium periodate is 1:0.45, and the reaction time is 5.5 h. The mass ratio of polydopamine-modified quaternized chitosan to oxidized hyaluronic acid is 1:1.4, the horseradish peroxidase concentration is 0.09 U / mL, the hydrogen peroxide concentration is 0.025%, and the in-situ gelation time is 11 min. The nano-silver content is 0.25% of the dry weight of the hydrogel. The hydrogel film thickness is 140 μm. Mesoporous silica nanoparticles have a particle size of 110 nm, a pore size of 3.2 nm, a specific surface area of ​​960 m² / g, a thiol density of 1.1 thiols / nm², and a polyethylene glycol molecular weight of 4000 Da. The mass ratio of ciprofloxacin hydrochloride to human defensin HD5 is 2.7:1, with a drug loading rate of 21%. Chitosan microspheres have a particle size of 12 μm, with honeysuckle glycoside added at 7% of the dry weight of the microspheres, resulting in an encapsulation efficiency of 88%. The polycaprolactone spinning solution concentration is 115 g / L, with drug-loaded mesoporous silica added at 4.5% of the mass of polycaprolactone and chitosan microspheres at 9.5%. The spinning voltage is 21 kV, the flow rate is 1.1 mL / h, and the receiving distance is 15 cm. The fiber diameter is 520 nm, the porosity is 76%, and the thickness is 110 μm. The hydrophobic fiber membrane has a polyvinylidene fluoride concentration of 195 g / L, a spinning voltage of 24 kV, a fiber diameter of 350 nm, a water contact angle of 152°, and a thickness of 55 μm. The gradient transition region has a thickness of 75 μm. The microchannels have a width of 110 μm, a depth of 65 μm, a spacing of 235 μm, and a microchannel density of 18 channels / cm². 2The hydrophilic fiber membrane has a polyvinyl alcohol concentration of 115 g / L, a degree of polymerization of 1900, a fiber diameter of 270 nm, and a thickness of 65 μm. Sodium carboxymethyl cellulose has a degree of substitution of 0.87 and a total solids content of 85 g / L. N,N'-methylenebisacrylamide is used at 0.7% of the total monomer mass. The porosity is 96%, the average pore size is 125 μm, and the dry thickness is 2.8 mm. The polyurethane aqueous dispersion has a solids content of 7% and a coating weight of 18 g / m². Hot-pressing temperature is 72℃, pressure is 0.45 MPa, and time is 55 s. The irradiation dose is 28 kGy. The resulting dressing has a total thickness of 4.2 mm and a basis weight of 210 g / m².

[0071] Comparative Example 1: Perianal medicated dressing without dopamine modification

[0072] The preparation method of this comparative example is basically the same as that of Example 2, except that the wet tissue adhesion layer is not modified with dopamine; instead, a hydrogel film is prepared directly using quaternized chitosan and oxidized hyaluronic acid. Specifically, the quaternized chitosan solution and the oxidized hyaluronic acid solution are mixed at a mass ratio of 1:1.35, without adding dopamine hydrochloride, and gelled in situ directly through a Schiff base reaction. Due to the lack of dopamine's catechol groups, this hydrogel film mainly relies on hydrogen bonding and electrostatic interactions to bind to the tissue surface, resulting in a significantly reduced adhesion strength. The preparation methods of the other layers are exactly the same as in Example 2. This comparative example is used to verify the key role of polydopamine modification in achieving chemical adhesion of wet tissues.

[0073] Comparative Example 2: Perianal medicated dressings without a dual-carrier drug sustained-release system

[0074] The preparation method of this comparative example is basically the same as that of Example 2, except that the drug sustained-release layer only uses a simple drug impregnation method, without using mesoporous silica nanoparticles and chitosan microsphere carrier systems. Specifically, ciprofloxacin hydrochloride, human defensin HD5, recombinant human epidermal growth factor, and honeysuckle glycoside are directly dissolved in polycaprolactone spinning solution according to the proportions in Example 2, and drug-loaded fiber membranes are prepared by electrospinning. Due to the lack of carrier protection and controlled release, the drug may be inactivated during the spinning process, and the release behavior exhibits the characteristics of initial burst release and insufficient release in the later stage, making it impossible to achieve time-controlled release. The preparation methods of other layers are exactly the same as in Example 2. This comparative example is used to verify the importance of a dual-carrier time-release system for achieving precise drug delivery.

[0075] Comparative Example 3: Perianal medicated dressing without a directional flow layer

[0076] The preparation method of this comparative example is basically the same as that of Example 2, except that the directional drainage layer is omitted, and the drug sustained-release layer is directly compounded with the highly absorbent outer layer. Due to the lack of a gradient wettable fiber membrane and a biomimetic microchannel array, the dressing relies solely on passive diffusion and capillary action for exudate transport, resulting in low transport efficiency. This easily leads to exudate accumulation on the wound surface causing wound maceration, or lateral diffusion of exudate causing dressing leakage. The preparation methods of other layers are exactly the same as in Example 2. This comparative example is used to verify the crucial role of the directional drainage layer in achieving efficient exudate management.

[0077] Comparative Example 4: Perianal medicated dressing without bionic microchannel array

[0078] The preparation method of this comparative example is basically the same as that of Example 2, except that although the directional flow guiding layer retains the sandwich-type gradient wettability fiber membrane structure, the biomimetic microchannel array prepared by femtosecond laser micromachining is omitted. Due to the lack of the additional capillary driving force and one-way valve effect provided by the microchannels, the directional transport efficiency of the exudate is reduced, and the one-way transport ratio is decreased. The preparation methods of other layers are exactly the same as those in Example 2. This comparative example is used to verify the unique role of the biomimetic microchannel array in enhancing the directional liquid transport capability.

[0079] Comparative Example 5: Three-layer perianal medicated dressing

[0080] This comparative example refers to the technical solution of Chinese Patent CN114571807A cited in the background section, and prepares a three-layer medical dressing. The first layer is a modified cotton fiber skin-friendly layer, which uses commercially available medical degreased cotton that has been chemically grafted and modified under alkaline conditions to introduce hydrophilic groups such as hydroxyethyl and carboxymethyl groups, thereby improving the fiber's moisture absorption and antibacterial properties. The second layer is a cross-linked polyimide fiber membrane, which is prepared by electrospinning and then thermal imidizing polyamic acid, mainly used to provide mechanical strength. The third layer is a polypropylene fiber protective layer, which serves as an isolation and protection function. The three layers are physically pressed together by simple hot-pressing. This dressing does not have chemical adhesion, drug delivery, or directional liquid transport functions, and is used for a comprehensive comparison with the technical solution of this invention.

[0081] To comprehensively evaluate the performance of the dressings of the present invention, a series of in vitro tests and animal experiments were conducted. The test samples included the dressings prepared in Examples 1 to 5, the dressings prepared in Comparative Examples 1 to 5, and commercially available perianal wound dressings as controls.

[0082] Test 1: Wet tissue adhesion strength test

[0083] The wet tissue adhesion properties of the dressings were quantitatively evaluated using a texture analyzer. Fresh, detached pigskin was used as the test substrate. After cleaning, the pigskin was soaked in physiological saline for 30 minutes to simulate a moist wound surface. Dressing samples were cut into 2cm × 2cm squares, attached to the pigskin surface, and a pre-pressure of 200g was applied for 30 seconds to ensure adequate contact. A vertical tensile peel test was then performed at a speed of 1mm / s, and the maximum adhesion force and adhesion work were recorded. Each sample was tested 10 times, and the average value was taken.

[0084] Test results showed that the adhesion strengths of the dressings prepared in Examples 1 to 5 on the surface of moist pigskin were 25.2 kPa, 30.5 kPa, 34.8 kPa, 32.1 kPa, and 31.7 kPa, respectively, all significantly higher than the 12.3 kPa of Comparative Example 1 and the 8.7 kPa of Comparative Example 5. Comparative Examples 2 to 4, due to the retention of the polydopamine-modified adhesive layer, had adhesion strengths close to those of the corresponding examples, at 29.8 kPa, 30.1 kPa, and 30.3 kPa, respectively. The adhesion strength of the commercial dressing was 10.5 kPa. These data indicate that polydopamine modification is crucial for achieving strong adhesion to moist tissue. The catechol groups of polydopamine can undergo various chemical bonding reactions with the amino, thiol, and hydroxyl groups on the tissue surface, including Schiff base reactions, Michael addition reactions, and hydrogen bonding interactions. This synergistic effect of multiple interactions significantly improves the adhesion strength. Comparative Example 1, lacking dopamine modification, relied solely on the electrostatic interactions and hydrogen bonding of quaternized chitosan, resulting in an adhesion strength only about 40% of that of the Examples. Example 3, with the highest dopamine grafting and the introduction of an enzyme-catalyzed cross-linked double cross-linking network, exhibited the highest adhesion strength of 34.8 kPa. The performance of the dressing under repeated adhesion conditions was further tested, simulating a re-adhesion scenario after cleaning a perianal wound. After peeling the dressing from the pigskin surface, the pigskin surface was gently rinsed with PBS buffer, and then the same dressing was reattached to test the adhesion strength. The results showed that the dressings of Examples 2 to 4 maintained 78% to 85% of the initial adhesion strength upon second adhesion, indicating that the adhesion mechanism of polydopamine has good reversibility and repeatability. This characteristic is particularly important for the clinical application of perianal wounds, as the dressing can be firmly re-adheded after defecation and cleaning, eliminating the need for replacement.

[0085] Test 2: Drug Release Kinetics Study

[0086] The drug release behavior in dressings was studied using ultraviolet-visible spectrophotometry and enzyme-linked immunosorbent assay (ELISA). A 1cm × 1cm dressing sample was immersed in 50mL of simulated wound solution, which consisted of pH 6.0 PBS buffer containing 1 mg / mL lysozyme, 10 mmol / L glutathione, and 10 mg / mL bovine serum albumin. The solution was placed in a 37℃ incubator and shaken at 100 rpm. Samples were taken at 0.5h, 1h, 2h, 4h, 8h, 12h, 24h, and on days 3, 5, 7, 10, and 14, with 2mL samples taken each time and an equal volume of fresh simulated wound solution added. The concentration of ciprofloxacin hydrochloride was determined by ultraviolet spectrophotometry at 276nm, and the concentration was calculated based on a standard curve. The concentrations of human defensin HD5 and recombinant human epidermal growth factor were determined using the corresponding ELISA kits. The concentration of honeysuckle glycosides was determined by high-performance liquid chromatography (HPLC).

[0087] Test results showed that the dressings prepared in Examples 1 to 5 exhibited ideal biphasic release curves. Ciprofloxacin hydrochloride was rapidly released within the first 4 hours, with a cumulative release of 40% to 50% of the total load over 4 hours and 70% to 75% over 24 hours. This rapid release pattern enabled the effective antibacterial concentration to be reached quickly during the acute phase of infection. The release curve of human defensin HD5 was similar to that of ciprofloxacin hydrochloride but slightly delayed, with a cumulative release of 65% to 70% over 24 hours. The synergistic release of the two antibacterial components enhanced the anti-infective effect. In contrast, recombinant human epidermal growth factor and honeysuckle glycosides exhibited a slow and sustained release characteristic. The release was slower in the first 3 days, with a cumulative release of approximately 20% to 25% over 3 days. Subsequently, the release rate gradually increased, reaching a cumulative release of 45% to 50% over 7 days and exceeding 60% over 14 days. This slow release pattern ensured a continuous supply of growth factors during the tissue remodeling phase. In Comparative Example 2, due to the lack of a carrier system for controlled release, ciprofloxacin hydrochloride experienced a burst release of over 80% within the first 2 hours and was almost completely released after 24 hours. In contrast, recombinant human epidermal growth factor, due to exposure to organic solvents and a high electric field during the spinning process, exhibited significantly reduced activity, with the actual effective release amount being less than 30% of the theoretical loading. This comparison fully demonstrates the superiority of the dual-carrier sequential release system.

[0088] Further analysis of the drug release mechanism revealed that the rapid release of ciprofloxacin hydrochloride was primarily due to the redox-responsive cleavage of polyethylene glycol (PEG) segments on the surface of mesoporous silica nanoparticles. In a simulated wound solution containing 10 mmol / L glutathione, disulfide bonds were rapidly reduced and cleaved, causing PEG segments to detach and opening pathways for drug diffusion. In the control experiment without glutathione, drug release was significantly slowed, with a cumulative release of only about 45% over 24 hours, demonstrating the effectiveness of the responsive release mechanism. The slow release of recombinant human epidermal growth factor (rGF) was mainly controlled by the enzymatic degradation process of chitosan microspheres. Lysozyme gradually degraded the cross-linked network of the microspheres by cleaving β-1,4-glycosidic bonds in the chitosan molecular chain, gradually releasing the encapsulated protein drug. In the control experiment without lysozyme, the cumulative release of rGF within 14 days was less than 20%, indicating that enzymatic degradation was the key factor controlling release. As a small molecule compound, honeysuckle glycosides have a release rate between that of antibiotics and growth factors, primarily influenced by the diffusion resistance and degradation rate of chitosan microspheres. This ingeniously designed dual-carrier system achieves the time-sequential release of antibacterial drugs and healing-promoting factors, perfectly matching the needs of different stages of wound healing.

[0089] Test 3: Evaluation of Unidirectional Liquid Transfer Performance

[0090] A self-made vertical liquid transfer testing device was used to quantitatively evaluate the directional flow capability of dressings. The device consisted of two transparent acrylic reservoirs, one above the other, with a dressing sample held in between. Simulated wound exudate was added to the upper reservoir. The simulated exudate was an aqueous solution containing 30 g / L bovine serum albumin, 9 g / L sodium chloride, and red dye, with a viscosity adjusted to 2-4 mPa·s to simulate actual wound exudate. The increase in liquid mass in the lower reservoir was monitored in real time using a high-precision electronic balance, and the liquid transfer rate was calculated. Simultaneously, a reverse test was performed, adding liquid from the lower reservoir to observe whether backflow occurred.

[0091] Test results show that the dressings prepared in Examples 1 to 5 exhibit excellent unidirectional liquid transport performance. The forward transport rates were 2.5 cm / s, 3.2 cm / s, 3.8 cm / s, 3.5 cm / s, and 3.3 cm / s, respectively, and the reverse transport rates were 0.30 cm / s, 0.35 cm / s, 0.42 cm / s, 0.38 cm / s, and 0.36 cm / s, respectively, with unidirectional transport ratios of 8.3, 9.1, 9.0, 9.2, and 9.2, respectively. In Comparative Example 3, due to the omission of the directional flow layer, the forward transport rate decreased to 1.2 cm / s, the reverse transport rate was 0.8 cm / s, and the unidirectional transport ratio was only 1.5, indicating that the liquid transport was basically bidirectional symmetrical diffusion. Comparative Example 4, while retaining the gradient wettability fiber membrane, lacks a biomimetic microchannel array. Its forward transport rate is 2.1 cm / s, its reverse transport rate is 0.5 cm / s, and its unidirectional transport ratio is 4.2, placing its performance between that of Example 1 and Comparative Example 3. Comparative Example 5 employs a three-layer homogeneous structure, with a forward transport rate of only 0.8 cm / s, a reverse transport rate of 0.6 cm / s, and a unidirectional transport ratio of 1.3, exhibiting almost no directional transport capability. The commercial dressing has a forward transport rate of 1.5 cm / s, a reverse transport rate of 0.7 cm / s, and a unidirectional transport ratio of 2.1.

[0092] These data fully demonstrate the rationality and effectiveness of the directional flow guiding layer design of this invention. The hydrophobic-hydrophilic gradient wettability fiber membrane structure provides a continuous Laplace pressure driving force. According to the Laplace pressure formula ΔP=γ(cosθ2-cosθ1) / r, the gradient flow guiding layer of Example 2 can generate a pressure difference of approximately 300Pa to 500Pa under the conditions of a water contact angle of 150° in the hydrophobic region, a water contact angle of less than 15° in the hydrophilic region, and an average pore size of approximately 500nm between fibers. This pressure difference is sufficient to drive the rapid transport of permeate with a viscosity of 2 to 4 mPa·s. The introduction of the biomimetic microchannel array further enhances the directional transport effect. The asymmetric wedge cross-section structure makes the capillary force of the liquid in the downstream direction much greater than that in the upstream direction. When the downstream sidewall angle is 30°, the capillary rise height h=2γcosθ / (ρgr) can reach a large value, while when the upstream sidewall angle is 60°, the capillary rise height is significantly reduced, forming a one-way valve effect. Comparing the data from Comparative Example 3 and Comparative Example 4 reveals that both the gradient wettability fiber membrane and the biomimetic microchannel array are indispensable for achieving efficient directional transport, and their synergistic effect is necessary to achieve the best results.

[0093] Furthermore, the dynamic transport process of droplets on a gradient surface was captured using a high-speed camera. It was observed that after contact with the gradient surface, the droplets rapidly spread and transported along the direction of increasing wettability, with a transport speed much faster than diffusion on a homogeneous surface. Liquid transport within the microchannels exhibited a clear directionality, with liquid preferentially transported along the long axis of the microchannel, while lateral diffusion was significantly suppressed. This anisotropic liquid transport behavior is key to achieving efficient percolation and diversion.

[0094] Test 4: Liquid Absorption and Retention Performance Test

[0095] Cut the dressing sample into 5cm × 5cm cubes, accurately weigh the initial mass m0, and then completely immerse the sample in physiological saline or simulated wound exudate. After a specified time, remove the sample, gently absorb the surface liquid with filter paper, weigh the mass m1 after absorption, and calculate the absorption ratio Q = (m1 - m0) / m0. In the liquid retention performance test, place the fully absorbed sample on filter paper, apply a pressure of 20 kPa for 2 minutes, and then weigh the mass m2. Calculate the liquid retention rate R = (m2 - m0) / (m1 - m0) × 100%.

[0096] Test results show that the absorbency ratios of the dressings prepared in Examples 1 to 5 in physiological saline were 155 g / g, 180 g / g, 190 g / g, 185 g / g, and 187 g / g, respectively, and the absorbency ratios in simulated wound exudate were 82 g / g, 95 g / g, 102 g / g, 98 g / g, and 100 g / g, respectively. The fluid retention rates were 85%, 89%, 91%, 90%, and 90%, respectively. Comparative Examples 1 to 4, due to retaining the highly absorbent outer layer, had absorbency performance close to that of the corresponding examples. The absorbency ratio of Comparative Example 5 was 45 g / g in physiological saline and 28 g / g in simulated wound exudate, with a fluid retention rate of 70%, significantly lower than that of the dressing of the present invention. The absorbency ratio of the commercial dressing was 120 g / g in physiological saline and 75 g / g in simulated wound exudate, with a fluid retention rate of 80%.

[0097] These data demonstrate that the high-absorption outer layer of this invention possesses superior liquid absorption capacity and excellent liquid retention performance. The high liquid absorption ratio of the sodium carboxymethyl cellulose and sodium polyacrylate composite sponge is mainly attributed to the Donnan osmotic pressure generated by the high-density negatively charged network and the three-dimensional porous structure formed by freeze-drying. The carboxyl groups on the sodium carboxymethyl cellulose and sodium polyacrylate molecular chains ionize in aqueous solution, generating negative charges. To maintain electroneutrality, a large number of counterions and water molecules enter the interior of the polymer network, leading to extreme swelling of the network. Simultaneously, the cross-linking effect of N,N'-methylenebisacrylamide limits the excessive extension of the polymer chains, achieving a balance between swelling pressure and the elastic recovery force of the cross-linked network. The interconnected porous structure formed during freeze-drying not only provides a large liquid storage space but also provides channels for rapid liquid permeation. Example 3, due to the use of the highest solid content and cross-linking agent dosage, resulted in the densest cross-linked network, exhibiting the highest liquid absorption ratio and liquid retention rate. Simulated wound exudate contains macromolecules such as proteins and cell debris, which can clog pores and increase solution viscosity to some extent. Therefore, the absorption rate is lower than that of physiological saline, but still remains at a high level of over 80 g / g. Excellent fluid retention ensures that the absorbed fluid will not leak out due to external pressure or dressing movement, which is crucial for preventing dressing leakage and maceration of the skin around the wound.

[0098] Test 5: Evaluation of Antibacterial Performance

[0099] The antibacterial properties of the dressings were quantitatively evaluated using the shaking flask method. Test strains included Staphylococcus aureus ATCC 25923, Escherichia coli ATCC 25922, and Pseudomonas aeruginosa ATCC 27853, all common pathogens causing wound infections in clinical practice. The strains were inoculated into nutrient broth and cultured to the logarithmic growth phase, with the bacterial concentration adjusted to 1×10⁻⁶. 6 CFU / mL. A 1cm × 1cm dressing sample was placed in an Erlenmeyer flask containing 10mL of bacterial suspension and incubated with shaking at 37℃ and 120rpm. Samples were taken at 0h, 1h, 3h, 6h, 12h, and 24h after contact, serially diluted, and plated to count the number of surviving colonies. The antibacterial rate was calculated as: (Control group colony count - Experimental group colony count) / Control group colony count × 100%.

[0100] Test results showed that the dressings prepared in Examples 1 to 5 exhibited strong antibacterial activity against all three test strains. Taking Staphylococcus aureus as an example, the antibacterial rate reached 85% to 90% after 1 hour of contact, 95% to 98% after 6 hours, and exceeded 99.5% after 24 hours. The antibacterial effect against Escherichia coli and Pseudomonas aeruginosa was slightly lower than against Staphylococcus aureus, but the antibacterial rate still reached over 99% after 24 hours. Comparative Example 1, lacking dopamine modification, failed to form in-situ silver nanoparticles, achieving an antibacterial rate of 92% to 94% after 24 hours, mainly relying on the contact antibacterial effect of quaternized chitosan and the efficacy of sustained-release antibiotics. Comparative Example 2, lacking a drug carrier system, experienced rapid antibiotic release in the first few hours, achieving an antibacterial rate of over 90% after 1 hour, but due to insufficient drug release in the later stages, the antibacterial rate dropped below 75% after 7 days. The antibacterial performance of Comparative Examples 3 to 5 was similar to that of the corresponding examples retaining the antibacterial layer. Commercial dressings do not contain antibacterial ingredients, and their 24-hour antibacterial rate is only about 35%.

[0101] Further testing of the dressing's antibacterial durability was conducted by immersing dressing samples in PBS buffer for 7 days, with the buffer changed daily to simulate the dilution and flushing effect of wound exudate on the drug. After 7 days, the samples were removed for antibacterial testing. The 24-hour antibacterial rate in Examples 1 to 5 remained between 95% and 97%, indicating that the dressing possesses long-lasting antibacterial activity. This sustained antibacterial effect is attributed to the synergistic effect of multiple antibacterial mechanisms: the positive charge of quaternized chitosan provides the first line of defense through electrostatic adsorption and disruption of bacterial cell membranes; the continuous and slow release of silver ions from the nanoparticles binds to thiol groups on the bacterial cell membrane, disrupting the bacterial respiratory chain and DNA replication; the sustained release of ciprofloxacin hydrochloride and human defensin HD5 from mesoporous silica nanoparticles provides potent drug bactericidal action; ciprofloxacin hydrochloride inhibits bacterial DNA gyrase, and defensin HD5 disrupts bacterial cell membrane integrity, with both synergistic effects; the antibacterial and anti-inflammatory effects of honeysuckle glycosides further enhance the overall anti-infective capacity.

[0102] A bacterial biofilm inhibition experiment was also conducted to evaluate the dressing's inhibitory effect on bacterial biofilm formation. The dressing samples were co-cultured with bacteria for 72 hours, and the amount of biofilm formed was quantitatively analyzed using crystal violet staining. The results showed that the biofilm inhibition rate of Examples 1 to 5 reached 82% to 88%, while the control group formed a large amount of dense biofilm. Bacterial biofilm is a significant cause of chronic wound healing difficulties; bacteria within the biofilm can increase antibiotic resistance by 100 to 1000 times. The dressing of this invention can effectively inhibit biofilm formation, which is of great significance for preventing the chronicity of perianal wounds.

[0103] Test 6: Cell Compatibility Evaluation

[0104] The cytocompatibility of the dressings was assessed using immortalized human keratinocytes (HaCaT) and human umbilical vein endothelial cells (HUVEC). Dressing samples were sterilized with 75% ethanol and then immersed in cell culture medium for 24 h at a volume ratio of 1:10 to obtain the dressing extract. Cells were seeded in 96-well plates at a density of 5 × 10³ cells / well and cultured for 24 h until cell attachment. The wells were then replaced with the dressing extract and cultured for another 24 h. Cell proliferation activity was assessed using the CCK-8 assay after 1, 3, 5, and 7 days of culture. Simultaneously, a live / dead cell staining assay was performed using calcein-AM and propidium iodide double staining, and cell viability was observed under a fluorescence microscope.

[0105] Test results showed that the dressing extracts prepared in Examples 1 to 5 did not exhibit significant cytotoxicity to either cell type. After 1 day of culture, the relative proliferation rate of HaCaT cells in the dressing extract was 95% to 102%, and after 7 days of culture, the relative proliferation rate was 103% to 115%, indicating that the dressing was not only non-toxic but also promoted the proliferation of keratinocytes. The proliferation of HUVEC cells in the dressing extract was similar to that of HaCaT cells, with a relative proliferation rate of 105% to 118% after 7 days of culture. Live and dead cell staining results showed that the cell viability was greater than 95% after 7 days of culture, with normal cell morphology and good extension. The cell compatibility of Comparative Examples 1 to 4 was close to that of the corresponding examples. Comparative Example 5, due to the use of a polyimide fiber layer, contained trace amounts of unreacted polyamic acid and small molecule byproducts generated during the imidization process in its extract, exhibiting slight toxicity to cells, with a relative cell proliferation rate of 85% to 88% after 7 days of culture.

[0106] Further cell adhesion experiments were conducted. The dressing sample was cut into 1cm × 1cm cubes and placed in 24-well plates, where HaCaT cells were seeded at a density of 2 × 10⁶ cells / well. 4 Cells / well. After culturing for 2 h, 4 h, 8 h, and 24 h, unadhered cells were gently washed with PBS buffer to remove them, and the number of adherent cells was quantitatively analyzed using crystal violet staining. The results showed that Examples 2 to 5, due to the presence of hyaluronic acid in the adhesion layer, exhibited a promoting effect on cell adhesion. After 24 h of culture, the number of adherent cells increased by 40% to 60% compared to the control group. Hyaluronic acid is an important component of the extracellular matrix and can bind to CD44 receptors on the cell surface, promoting cell adhesion, migration, and proliferation. This property is beneficial for epithelial cell colonization and migration on the wound surface, accelerating the reepithelialization process of the wound.

[0107] Test 7: Evaluation of in vitro healing-promoting activity

[0108] The effects of dressing extract on cell migration were assessed using cell scratch assays and Transwell migration assays, and the regulatory effect of the extract on cell secretion of healing-promoting factors was detected using ELISA. HaCaT cells were cultured to confluence, and scratches were made on a monolayer of cells using a 200 μL pipette tip. Exfoliated cells were gently washed away with PBS buffer. The culture medium was then replaced with dressing extract, and the cells were cultured for another period. Scratch widths were recorded at 0 h, 12 h, and 24 h, and the scratch healing rate was calculated as (0 h scratch width - 24 h scratch width) / 0 h scratch width × 100%. In the Transwell migration assay, cells were seeded in the upper chamber of a Transwell apparatus, and dressing extract was added to the lower chamber as a chemokine. After 24 h of culture, cells were fixed and stained, and the number of cells migrating to the lower chamber was counted.

[0109] Test results showed that the dressing extracts from Examples 1 to 5 significantly promoted cell migration. The 24-hour scratch healing rate was 75% to 85%, while the control group was only about 45%. In the Transwell assay, the number of migrating cells increased by 60% to 80% compared to the control group. This cell migration-promoting effect was mainly attributed to the sustained-release recombinant human epidermal growth factor (EGFR) and honeysuckle glycosides in the dressing. EGFR binds to the EGFR receptor on the cell surface, activating downstream MAPK and PI3K / Akt signaling pathways, promoting cell proliferation, migration, and differentiation. Honeysuckle glycosides create a favorable microenvironment for cell migration by inhibiting the expression of inflammatory factors. In Comparative Example 2, due to the lack of a drug delivery system, EGFR was largely inactivated during preparation, significantly weakening its cell migration-promoting effect; the 24-hour scratch healing rate was only about 55%.

[0110] ELISA results showed that the dressing extracts from Examples 1 to 5 upregulated type I and type III collagen secreted by fibroblasts. After 7 days of culture, type I collagen secretion increased by 50% to 70% compared to the control group, and type III collagen secretion increased by 45% to 65%. Collagen is a major component of newly formed granulation tissue during wound healing, and its increased synthesis is beneficial for restoring the mechanical strength of the wound. Simultaneously, the dressing extracts upregulated the expression of vascular endothelial growth factor (VEGF), with VEGF secretion increasing by 55% to 75% after 7 days of culture compared to the control group. VEGF is a key factor in promoting angiogenesis, which is crucial for providing oxygen and nutrients to newly formed tissue. These in vitro experimental results demonstrate that the dressing of this invention can not only control infection but also actively promote the wound healing process through the sustained release of bioactive factors.

[0111] The following table summarizes the results of the key test metrics:

[0112] Wet tissue adhesion strength (kPa) 25.2 30.5 34.8 32.1 31.7 12.3 29.8 30.1 30.3 8.7 10.5 24-hour antibiotic release rate (%) 70 72 75 73 72 71 85 72 72 - - 14-day growth factor release rate (%) 60 63 67 65 64 62 28 63 63 - - Forward liquid transport rate (cm / s) 2.5 3.2 3.8 3.5 3.3 3.1 3.2 1.2 2.1 0.8 1.5 One-way transmission ratio 8.3 9.1 9 9.2 9.2 9 9.1 1.5 4.2 1.3 2.1 Absorption rate - physiological saline (g / g) 155 180 190 185 187 178 180 180 180 45 120 Liquid absorption ratio - simulated exudate (g / g) 82 95 102 98 100 93 95 95 95 28 75 Liquid retention rate (%) 85 89 91 90 90 88 89 89 89 70 80 24h Antibacterial Rate - Staphylococcus aureus (%) 99.5 99.7 99.8 99.7 99.7 93.5 99.6 99.7 99.7 35 35 24h antibacterial rate - Escherichia coli (%) 99.2 99.5 99.6 99.5 99.5 92.8 99.4 99.5 99.5 32 33 Antibacterial rate (%) after 7 days 95 96 97 96 96 88 72 96 96 30 15 Relative cell proliferation rate - 7 days (%) 103 108 115 110 112 105 106 108 108 86 98 Cell viability (%) 95 96 97 96 96 95 96 96 96 88 94 24-hour scratch healing rate (%) 75 80 85 82 81 78 55 80 80 45 48

[0113] Test 8: Evaluation of Animal Experiments

[0114] To comprehensively evaluate the in vivo healing-promoting effect and biosafety of the dressing, animal experiments were conducted using a rat full-thickness skin defect model and a rabbit perianal infected wound model. The animal experimental protocols were reviewed and approved by the institution's animal ethics committee, and the experimental procedures strictly adhered to animal welfare and ethical principles.

[0115] A full-thickness skin defect model was established using adult male SD rats weighing 200-220g. After anesthetizing the rats with sodium pentobarbital via intraperitoneal injection, the backs were shaved and disinfected. A circular full-thickness skin defect with a diameter of 15mm, reaching the fascia layer, was prepared using a punch. The animals were randomly divided into 6 groups of 10 each: a blank control group, a commercial dressing control group, Example 2 group, Example 3 group, Comparative Example 2 group, and Comparative Example 3 group. The corresponding dressings were cut into 20mm diameter circles and applied to the wound, secured with breathable medical tape. The dressings were changed every 3 days, and the wound appearance was photographed and the wound area measured. Five animals were sacrificed on days 7 and 14, and wound tissue was collected for histological analysis.

[0116] Wound healing results showed that the healing effects of groups 2 and 3 were significantly better than other groups. The wound healing rates on day 7 were 65% and 70%, respectively, and on day 14, they were 88% and 92%, respectively. In the blank control group, the wound healing rates on days 7 and 14 were 42% and 68%, respectively, while in the commercial dressing control group, they were 50% and 75%, respectively. In comparative group 2, due to improper drug release, the healing rate on day 7 was 55%, and on day 14, it was 78%. In comparative group 3, due to improper exudate management, some animals experienced wound maceration and surrounding skin damage; the healing rate on day 7 was 48%, and on day 14, it was 72%. Histological analysis showed that on day 7, the wound bed in groups 2 and 3 had formed abundant granulation tissue, densely distributed neovascularization, and mild inflammatory cell infiltration. By day 14, the wound surface had completed re-epithelialization, the epidermal structure was intact, the collagen fibers in the dermis were arranged in an orderly manner, and accessory organs began to regenerate. Immunohistochemical staining for CD31 was used to assess angiogenesis density. The angiogenesis density in groups 2 and 3 was increased by 62% and 68% respectively compared to the control group, indicating that the growth factors released by the dressing effectively promoted angiogenesis. Masson trichrome staining was used to assess collagen deposition. The collagen fibers in the example groups were more orderly arranged, and the ratio of type I collagen to type III collagen was closer to that of normal skin, suggesting higher quality wound remodeling.

[0117] The rabbit perianal infection wound model used adult New Zealand white rabbits weighing 2.5 to 3.0 kg. After anesthetizing the animals with ketamine via intramuscular injection, two 1.5 cm × 1.5 cm full-thickness skin defects were prepared at the 3 o'clock and 9 o'clock positions around the anus. The wounds were inoculated with a mixed bacterial solution containing 1 × 10⁻⁶ Staphylococcus aureus and 1 × 10⁻⁶ Escherichia coli. 7 An infected wound model was established using CFU / mL. Dressings were applied 24 hours later. Animals were randomly divided into four groups of six each: a commercial dressing control group, Example 2 group, Example 3 group, and Comparative Example 5 group. Dressings were changed every three days to observe adhesion, exudate management, and wound healing progress. Wound secretions were collected on days 7 and 14 for bacterial culture and counting. Animals were sacrificed on day 21 for wound tissue histological analysis.

[0118] Experimental results showed that the dressings in Examples 2 and 3 exhibited excellent adhesion properties in the perianal area of ​​rabbits. The dressings remained stably attached during defecation and daily activities, and even after cleaning, they could be firmly reattached with gentle pressing, requiring no replacement within 3 days. The dressings in the commercial dressing control group and Comparative Example 5 showed poor adhesion properties; most animals experienced partial dressing detachment after defecation, requiring daily dressing changes. Wound bacterial counts showed that on day 3, the bacterial count in Examples 2 and 3 groups was reduced by more than 99% compared to the initial inoculation amount, and by day 7, there was virtually no bacterial growth in the wound. In the commercial dressing control group, the bacterial count decreased by approximately 70% on day 3, but a significant amount of bacteria remained on day 7. Comparative Example 5 showed the worst antibacterial effect, with only a 50% reduction in bacterial count on day 3, and it remained at a high level on day 7. The potent antibacterial effect of the Example groups effectively prevented secondary infection and biofilm formation in the wound. Regarding exudate management, the dressings in Examples 2 and 3 effectively absorbed and drained exudate, keeping the outer surface of the dressings dry throughout, with no side leakage observed. In Comparative Example 5, due to the lack of an efficient exudate management system, some animals experienced dressing saturation and side leakage on days 3 to 5, leading to perianal skin maceration and secondary dermatitis. Wound healing assessment on day 21 showed that the complete healing rates in Examples 2 and 3 were 83% and 92%, respectively. The healed skin color and texture were close to normal tissue, with minimal scarring. The complete healing rate in the commercial dressing control group was 67%, while in Comparative Example 5 it was only 50%, with some animals developing chronic, difficult-to-heal wounds due to persistent infection and exudate irritation. Histological analysis confirmed that the healed tissue structure in the Example groups was closer to normal skin, with moderate epidermal thickness, regular arrangement of collagen fibers in the dermis, complete resolution of inflammation, and no significant scar hyperplasia.

[0119] Experimental results from a rabbit perianal infection wound model fully demonstrate the superior performance of the dressing of this invention in practical clinical applications. The dressing's strong adhesion solves the problem of easy dressing detachment in the perianal area; its efficient exudate management system prevents wound maceration and damage to surrounding skin; its long-lasting antibacterial activity effectively controls infection; and its sustained-release bioactive factors promote tissue repair and regeneration. The synergistic effect of these functions shortens the healing time of perianal infection wounds by 30% to 40%, reducing the frequency of dressing changes from the traditional 1-2 times per day to once every 3 days, greatly improving treatment efficacy and patient comfort.

[0120] The above description is only a preferred embodiment of the present invention and is not intended to limit the present invention. Any modifications, equivalent substitutions and improvements made within the spirit and principles of the present invention should be included within the protection scope of the present invention.

Claims

1. A perianal medicated dressing made of multi-layered diversion composite fabric, characterized in that, From bottom to top, it includes a wet tissue adhesion layer, a drug sustained-release layer, a directional flow layer, and a highly absorbent outer layer, wherein: The wet tissue adhesion layer is made of a composite hydrogel membrane of polydopamine-modified quaternized chitosan and oxidized hyaluronic acid. The polydopamine-modified quaternized chitosan is obtained by grafting dopamine groups onto the quaternized chitosan molecular chain. The quaternized chitosan is prepared by reacting chitosan with a deacetylation degree of 85% to 95% with 3-chloro-2-hydroxypropyltrimethylammonium chloride at a mass ratio of 1:2 to 1:2.

5. The oxidized hyaluronic acid is prepared by oxidizing sodium hyaluronate with a molecular weight of 800kDa to 1200kDa with sodium periodate at a molar ratio of 1:0.3 to 1:0.

5. The polydopamine-modified quaternized chitosan and oxidized hyaluronic acid are mixed at a mass ratio of 1:1.2 to 1:1.5 and then gelled in situ through a Schiff base reaction to form a hydrogel membrane. The thickness of the hydrogel membrane is 100μm to 150μm. The drug sustained-release layer comprises a mesoporous silica nanoparticle drug-carrying system and a chitosan microsphere drug-carrying system. The mesoporous silica nanoparticle drug-carrying system loads ciprofloxacin hydrochloride and human defensin HD5, with a mass ratio of ciprofloxacin hydrochloride to human defensin HD5 of 2:1 to 3:

1. The mesoporous silica nanoparticles have a particle size of 80 nm to 120 nm, a pore size of 2.5 nm to 3.5 nm, and a specific surface area greater than 900 m² / g. The chitosan microsphere drug-carrying system encapsulates recombinant human epidermal cells. The product contains skin growth factor and honeysuckle glycoside, wherein the honeysuckle glycoside accounts for 5% to 8% of the dry weight of chitosan microspheres; the mesoporous silica nanoparticles and chitosan microspheres are dispersed in a polycaprolactone nanofiber membrane; the amount of mesoporous silica nanoparticles added is 3% to 5% of the mass of polycaprolactone; the amount of chitosan microspheres added is 8% to 10% of the mass of polycaprolactone; and the polycaprolactone nanofiber membrane has a fiber diameter of 300 nm to 600 nm, a porosity greater than 70%, and a thickness of 80 μm to 120 μm. The directional flow-guiding layer is a sandwich-type gradient wettable fiber membrane structure, comprising, from bottom to top, a hydrophobic region, a gradient transition region, and a hydrophilic region. The hydrophobic region is made of polyvinylidene fluoride nanofiber membrane with a fiber diameter of 200 nm to 400 nm, a water contact angle of 145° to 155°, and a thickness of 40 μm to 60 μm. The gradient transition region is made of a blend of polyvinylidene fluoride and polyvinylpyrrolidone nanofiber membrane with a thickness of 60 μm to 80 μm and a water contact angle from 1... The hydrophilic region, transitioning from 15° to 65°, is made of polyvinyl alcohol nanofiber membrane with a fiber diameter of 150nm to 300nm, a water contact angle of less than 15°, and a thickness of 50μm to 70μm. The surface of this gradient transition region is fabricated with a biomimetic microchannel array. These microchannels have an asymmetric wedge-shaped cross-section structure, a width of 80μm to 120μm, a depth of 50μm to 70μm, a spacing of 200μm to 250μm, and a microchannel density of 15 channels / cm². 2 Up to 20 strips / cm 2 ; The high-absorbency outer layer is made of a composite sponge of sodium carboxymethyl cellulose and sodium polyacrylate. The degree of substitution of sodium carboxymethyl cellulose is 0.8 to 0.9, the molecular weight of sodium polyacrylate is 2 million, the mass ratio of sodium carboxymethyl cellulose to sodium polyacrylate is 3:2, the porosity of the composite sponge is greater than 92%, the average pore size is 50 μm to 150 μm, and the dry thickness is 2 mm to 3 mm.

2. The perianal medicated dressing with multi-layered drainage composite fabric according to claim 1, characterized in that... The amount of dopamine grafted into the polydopamine-modified quaternized chitosan is 10% to 15% of the mass of the quaternized chitosan. The gel time of the hydrogel membrane at 37°C is 8 to 12 minutes. The adhesion strength of the hydrogel membrane to pigskin is 25 kPa to 35 kPa. The moisture content of the hydrogel membrane is 75% to 85%, and the oxygen permeability is greater than 800 mL / m²·24h.

3. The perianal medicated dressing with multi-layered drainage composite fabric according to claim 1, characterized in that... The surface of the mesoporous silica nanoparticles is modified with disulfide-bridged polyethylene glycol molecular chains. The thiol density on the surface of the mesoporous silica nanoparticles is 0.8 to 1.2 thiols / nm². The molecular weight of the polyethylene glycol is 2000 Da to 5000 Da. The loading rate of ciprofloxacin hydrochloride in the mesoporous silica nanoparticles is 18% to 22%, and 70% to 75% is released within 24 hours in a simulated wound solution at pH 6.0 containing 10 mmol / L glutathione.

4. The perianal medicated dressing with multi-layered drainage composite fabric according to claim 1, characterized in that... The chitosan microspheres have a particle size of 5 μm to 15 μm and are prepared by cross-linking with glutaraldehyde. The encapsulation rate of the recombinant human epidermal growth factor in the chitosan microspheres is greater than 85%, and the cumulative release amount in a simulated wound solution containing 1 mg / mL of lysozyme is greater than 60% within 14 days.

5. The perianal medicated dressing with multi-layered drainage composite fabric according to claim 1, characterized in that... The wet tissue adhesion layer further comprises silver nanoparticles, which are formed by in-situ reduction of silver nitrate with polydopamine. The content of the silver nanoparticles is 0.1% to 0.3% of the dry weight of the hydrogel. The wet tissue adhesion layer is further cured by cross-linking catalyzed by horseradish peroxidase, with the concentration of horseradish peroxidase being 0.05 U / mL to 0.1 U / mL and the concentration of hydrogen peroxide as an oxidant being 0.01% to 0.03%.

6. The perianal medicated dressing with multi-layered drainage composite fabric according to claim 1, characterized in that... The dressing layers are integrated through a hot-pressing composite process, with a polyurethane aqueous dispersion coated between the layers as an adhesive layer. The polyurethane aqueous dispersion has a solid content of 5% to 8%, a coating amount of 15 g / m² to 20 g / m², an interlayer peel strength greater than 15 N / 25 mm, and an interlayer gas permeability greater than 500 mL / m²·24 h. The total thickness of the dressing is 3.5 mm to 4.5 mm, and the basis weight is 180 g / m² to 220 g / m².

7. A method for preparing a perianal medicated dressing with a multilayer drainage composite fabric according to any one of claims 1 to 6, characterized in that... This includes the following steps: Step 1, Preparation of the wet tissue adhesion layer: Chitosan with a degree of deacetylation of 85% to 95% was prepared into an alkaline solution with a mass concentration of 2 g / L to 3 g / L, and the pH was adjusted to 9 to 10. 3-chloro-2-hydroxypropyltrimethylammonium chloride was added, with a mass ratio of chitosan to 3-chloro-2-hydroxypropyltrimethylammonium chloride of 1:2 to 1:2.

5. The mixture was stirred and reacted at 60°C to 70°C for 6 to 8 hours. After the reaction was completed, the mixture was purified by dialysis and freeze-dried to obtain quaternized chitosan. The quaternized chitosan was dissolved in PBS buffer at pH 8.5 to prepare a solution with a concentration of 15 g / L to 20 g / L. Dopamine hydrochloride was added. The amount added is 10% to 15% of the mass of quaternized chitosan. The mixture is reacted at room temperature in an oxygen atmosphere for 24 to 36 hours to obtain a polydopamine-modified quaternized chitosan solution. Sodium hyaluronate with a molecular weight of 800 kDa to 1200 kDa and sodium periodate are reacted at a molar ratio of 1:0.3 to 1:0.5 under light-protected conditions for 4 to 6 hours. After dialyzing, the mixture is freeze-dried to obtain oxidized hyaluronic acid. The polydopamine-modified quaternized chitosan solution and the oxidized hyaluronic acid solution are mixed at a mass ratio of 1:1.2 to 1:1.5 and in-situ gelled at 37°C via Schiff base reaction for 8 to 12 minutes to obtain a hydrogel film with a thickness of 100 μm to 150 μm. Step 2, Preparation of the drug sustained-release layer: Using hexadecyltrimethylammonium bromide as a template agent and tetraethyl orthosilicate as a silicon source, the mixture was prepared at a molar ratio of 1:8:140:9000 for hexadecyltrimethylammonium bromide, tetraethyl orthosilicate, ammonia, and water. The mixture was reacted at 80°C for 3 hours, and then calcined at 650°C to remove the template, yielding mesoporous silica nanoparticles. These nanoparticles were then silanized using 3-mercaptopropyltrimethoxysilane, and linked to a molecular weight of 2 through a disulfide crosslinking reaction. Polyethylene glycol segments ranging from 000 Da to 5000 Da were loaded onto mesoporous silica nanoparticles after ciprofloxacin hydrochloride and human defensin HD5 were mixed at a mass ratio of 2:1 to 3:

1. Chitosan with a degree of deacetylation greater than 85% was prepared into an acetic acid solution with a concentration of 25 g / L and emulsified with liquid paraffin containing 3% Span-80 emulsifier at a volume ratio of 1:

5. Glutaraldehyde crosslinking agent with a concentration of 5 g / L was added dropwise, and the reaction was carried out at 45°C for 2 h. After washing and drying, the desired product was obtained. Chitosan microspheres with a particle size of 5 μm to 15 μm were obtained. Recombinant human epidermal growth factor and honeysuckle glycosides were encapsulated in the chitosan microspheres. The amount of honeysuckle glycosides added was 5% to 8% of the dry weight of the chitosan microspheres. Polycaprolactone with a molecular weight of 80,000 Da was dissolved in a mixed solvent of chloroform and methanol in a volume ratio of 3:1 to prepare a spinning solution with a concentration of 100 g / L to 120 g / L. The drug-loaded mesoporous silica nanoparticles and chitosan microspheres were dispersed in the spinning solution. The amount of mesoporous silica nanoparticles added is 3% to 5% of the mass of polycaprolactone, and the amount of chitosan microspheres added is 8% to 10% of the mass of polycaprolactone. A nanofiber membrane is prepared using an electrospinning process with a spinning voltage of 18kV to 22kV, a flow rate of 0.8mL / h to 1.2mL / h, and a receiving distance of 15cm, resulting in a drug-release layer with a fiber diameter of 300nm to 600nm, a porosity greater than 70%, and a thickness of 80μm to 120μm. Step 3, Preparation of the Directional Guiding Layer: Polyvinylidene fluoride (PVDF) with a molecular weight of 300,000 was dissolved in a mixed solvent of N,N-dimethylformamide and acetone in a mass ratio of 6:4 to prepare a spinning solution with a concentration of 180 g / L to 200 g / L. Hydrophobic fiber membranes were prepared using electrospinning at a voltage of 20 kV to 25 kV, a flow rate of 1.0 mL / h, and a receiving distance of 18 cm. This yielded PVDF nanofiber membranes with fiber diameters of 200 nm to 400 nm, water contact angles of 145° to 155°, and thicknesses of 40 μm to 60 μm. A gradient transition region was prepared using coaxial electrospinning technology. The core layer consisted of a 150 g / L PVDF solution, and the shell layer consisted of a 120 g / L polyvinylpyrrolidone solution. Continuous wetting was achieved by varying the inner-outer layer flow rate ratio from 1:0.5 to 1:

2. A gradient was used to obtain blended fiber membranes with a thickness of 60 μm to 80 μm and a water contact angle that gradually changed from 115° to 65°. Polyvinyl alcohol with a degree of alcoholysis of 98% and a degree of polymerization of 1700 to 2000 was prepared into an aqueous solution with a concentration of 100 g / L to 120 g / L. Hydrophilic fiber membranes were prepared using electrospinning. After spinning, the fibers were crosslinked and cured with glutaraldehyde vapor at 25°C for 12 h to obtain polyvinyl alcohol nanofiber membranes with a fiber diameter of 150 nm to 300 nm, a water contact angle of less than 15°, and a thickness of 50 μm to 70 μm. A biomimetic microchannel array was fabricated on the surface of the gradient transition region using femtosecond laser micromachining technology. The microchannels had an asymmetric wedge-shaped cross-section structure with a width of 80 μm to 120 μm, a depth of 50 μm to 70 μm, a spacing of 200 μm to 250 μm, and a microchannel density of 15 channels / cm². 2 Up to 20 strips / cm 2 ; Step 4, Preparation of the high-absorption outer layer: Sodium carboxymethyl cellulose with a degree of substitution of 0.8 to 0.9 and sodium polyacrylate with a molecular weight of 2 million are mixed in an aqueous solution at a mass ratio of 3:2 to prepare a solution with a total solid content of 80 g / L to 100 g / L. 0.5% to 0.8% of N,N'-methylenebisacrylamide by mass of monomer is added as a crosslinking agent, 0.3% of ammonium persulfate is added as an initiator, and 0.2% of N,N,N',N'-tetramethylethylenediamine is added as a promoter. The solution is freeze-dried at -50°C and a vacuum degree of less than 10 Pa for 48 h to form a three-dimensional porous sponge structure with a porosity greater than 92%, an average pore size of 50 μm to 150 μm, and a dry thickness of 2 mm to 3 mm. Step 5, Interlayer Integration: The wet tissue adhesion layer, drug sustained-release layer, directional flow layer, and high-absorption outer layer are sequentially subjected to low-temperature oxygen plasma surface treatment. The plasma power is 80W and the treatment time is 30s. A polyurethane aqueous dispersion with a solid content of 5% to 8% is coated between each layer as an adhesive layer, with a coating amount of 15g / m² to 20g / m². Interlayer integration is carried out using a hot-pressing composite process. The composite temperature is 65℃ to 75℃, the pressure is 0.3MPa to 0.5MPa, and the time is 45s to 60s. A multilayer composite dressing with a total thickness of 3.5mm to 4.5mm and a basis weight of 180g / m² to 220g / m² is obtained. After being cut and shaped, it is sterilized by cobalt-60 gamma irradiation with an irradiation dose of 25kGy to 30kGy.

8. The preparation method according to claim 7, characterized in that... In step one, during the in-situ gelation of the wet tissue adhesion layer, horseradish peroxidase is added for enzymatic cross-linking. The concentration of horseradish peroxidase is 0.05 U / mL to 0.1 U / mL. Simultaneously, hydrogen peroxide with a concentration of 0.01% to 0.03% is added as an oxidant. The reaction is carried out at 37°C for 5 to 10 minutes to achieve a dual cross-linked network structure of Schiff base cross-linking and enzyme-catalyzed cross-linking. Silver nitrate solution is added to the polydopamine-modified quaternized chitosan solution to utilize the catechol groups of polydopamine to reduce silver ions in situ and form silver nanoparticles. The content of the silver nanoparticles is controlled to be 0.1% to 0.3% of the dry weight of the hydrogel.

9. The preparation method according to claim 7, characterized in that... In step four, the high-absorption outer layer undergoes secondary curing via visible light-initiated thiol-olefin click chemical crosslinking. Allyl and thiol functional groups are grafted onto the sodium carboxymethyl cellulose and sodium polyacrylate molecular chains, with a modification degree of 3 to 5 functional groups grafted per 100 monosaccharide units. Riboflavin or eosin Y is used as the photoinitiator, and the curing is carried out under a light intensity of 5 mW / cm². 2 Up to 10mW / cm 2 Crosslinking was completed under the condition of irradiation time of 3 to 5 minutes.

10. The use of the perianal medicated dressing of the multilayer drainage composite fabric according to any one of claims 1 to 6 in the preparation of a medical dressing for treating perianal infected wounds.