Diabetic foot infection targeting dressing and preparation method

Abstract

The application discloses a kind of diabetic foot infection targeting dressings and preparation method.The dressing includes sequentially stacked wound contact layer, gradient wetting transition layer and hydrophobic gas guide layer;Wound contact layer is the hydrophilic fiber layer of embedding pH response drug-loaded microspheres, microspheres are with silk glue protein load antibacterial drug as core, tannic acid-metal ion coordination polymer as shell layer;Gradient wetting transition layer has the multilayer fiber structure of decreasing wetting and the through cone-shaped pore;Hydrophobic gas guide layer is hydrophobic fiber layer.Through microsphere-fiber-pore three-stage size anchoring, coupling synergy of wetting gradient and pore size gradient, and negative pressure locking of hydrophobic gas guide layer, the time sequence decoupling of "one-way liquid flow-microsphere mechanical anchoring-drug environment response" is realized.Preparation method includes microsphere reverse emulsion method, multilayer sequential electrospinning and ultraviolet laser selective pore making.The dressing is used for diabetic foot infection targeted therapy, with one-way wetting, intelligent response drug release and wet healing function.

Description

Technical Field

[0001] This invention belongs to the field of medical biomaterials technology, specifically relating to an infection-targeting dressing for diabetic foot wound management and its preparation method. Background Technology

[0002] Diabetic foot ulcers are one of the most serious chronic complications of diabetes, characterized by difficult wound infection control and a long healing period, placing a heavy burden on patients and society. Exudate management in diabetic foot wounds is a crucial step in promoting healing: excessive exudate leads to wound maceration and damages newly formed tissue; insufficient exudate leads to wound dryness and hinders epithelial migration. Therefore, an ideal diabetic foot dressing should possess both effective exudate management and precise infection control functions.

[0003] In existing technologies, diabetic foot dressings mainly fall into two technical categories: The first category is responsive drug-release dressings, such as pH-responsive microspheres and enzyme-responsive hydrogels. These dressings can trigger drug release under the influence of the infection microenvironment (acidic pH or high enzyme activity), achieving targeted therapy for the infection. However, these dressings usually lack effective exudate drainage structures. Wound exudate continuously washes away the drug-loaded microspheres, causing the drug to be passively carried away by the liquid phase, severely disrupting the pulsatile drug release curve, reducing antibacterial efficacy, and increasing the risk of drug resistance.

[0004] The second category is unidirectional moisture-wicking dressings, such as Janus gradient fiber membranes and multilayer foam dressings. These dressings drive exudate outward from the wound side to the outside through a wettability gradient or pore size gradient, effectively preventing wound maceration. However, these dressings typically only have physical exudate management functions and lack the ability to intelligently respond to the infection microenvironment, failing to dynamically adjust drug release according to the wound infection status.

[0005] The two aforementioned technical approaches each address only part of the problem; no technical solution has yet achieved an organic integration of "responsive drug release" and "unidirectional moisture wicking" within the same thin-film unit. Those skilled in the art generally agree that the wound contact side requires hydrophilicity to absorb exudate and provide a moist healing environment, while the anchoring of drug-loaded microspheres requires a hydrophobic environment to prevent the drug from being washed away by exudate—a physical contradiction existing in the same spatial dimension. Simultaneously, responsive drug release requires the microspheres to be fully exposed to the wound environment (large fibrous pores), while unidirectional moisture wicking requires a dense gradient pore structure (small pores), creating a technical contradiction in the structural dimension. Therefore, there is a technical bias in the field regarding the integrated use of pH-responsive carriers and gradient fiber membranes, believing that the two are difficult to be compatible in the same dressing. Summary of the Invention

[0006] The present invention aims to overcome the above-mentioned defects in the prior art and provide a targeted dressing for diabetic foot infection and a preparation method thereof, which achieves the coexistence of hydrophilic moisture-wicking and hydrophobic drug-anchoring in the same thin film unit, without interfering with each other's functional logic.

[0007] To achieve the above objectives, the present invention adopts the following technical solution: A targeted dressing for diabetic foot infection includes a wound contact layer, a gradient wetting transition layer and a hydrophobic air-conducting layer stacked sequentially. The wound contact layer is a hydrophilic microsphere embedded in a fiber layer, consisting of a fiber network formed by interwoven hydrophilic polymer fibers and pH-responsive drug-loaded microspheres with a particle size larger than the pores of the fiber network and physically locked therein; the pH-responsive drug-loaded microspheres include: a core of porous sericin microspheres loaded with antibacterial drugs, and a tannic acid-metal ion coordination polymer shell coating the surface of the core. The gradient wetting transition layer is a multi-layer fiber structure in which the fiber wettability decreases from the wound contact layer to the hydrophobic air-conducting layer, and a conical channel is provided through the gradient wetting transition layer. The diameter of the conical channel gradually increases from the side closer to the wound contact layer to the side closer to the hydrophobic air-conducting layer. The hydrophobic air-conducting layer is a hydrophobic polymer fiber layer.

[0008] In this invention, the hydrophilic polymer fiber of the wound contact layer is preferably a co-spun fiber of polycaprolactone (PCL) and tannic acid-metal ion coordination polymer. Tannic acid provides a large number of phenolic hydroxyl groups, giving the fiber hydrophilicity, while iron ions impart photothermal-assisted antibacterial ability. The fiber diameter is preferably 200-500 nm, and the porosity is preferably 70%-85%, to ensure rapid exudate infiltration while providing sufficient anchoring space for the microspheres.

[0009] In this invention, the metal ions in the pH-responsive drug-loaded microspheres are preferably Fe. 3+ Cu 2+ The microspheres contain at least one of the following: the antibacterial drug is preferably selected from at least one of mupirocin and silver nanoparticles; the microsphere particle size is preferably 1-5 μm. The microspheres have a core-shell structure: the core is a porous sericin microsphere loaded with the antibacterial drug, and the shell is a tannic acid-metal ion coordination polymer. In an infected microenvironment (pH < 6.0), the coordination bonds in the shell break, the shell disintegrates, and the drug is released from the core; in a normal tissue environment (pH ≈ 7.4), the shell remains stable, and the drug is sealed. This particle size design ensures that the microspheres are larger than the interfiber gaps but do not block the fluid channels, achieving physical anchoring of the microspheres within the fiber network.

[0010] In this invention, the gradient wetting transition layer preferably comprises a first transition layer (polyacrylonitrile fiber, contact angle 30°-50°), a second transition layer (silk fibroin fiber, contact angle 50°-70°), and a third transition layer (polycaprolactone fiber, contact angle 70°-90°) stacked sequentially along the direction away from the wound contact layer. Simultaneously, the gradient wetting transition layer is permeated with conical channels. The pore size of the conical channels is preferably 3-5 μm on the side near the wound contact layer and 15-20 μm on the side near the hydrophobic air-conducting layer. The density of the conical pores is preferably 100-500 pores / cm². Based on the principle of wettability gradient driving, since the contact angle of the fibers in the gradient wetting transition layer gradually increases along the direction away from the wound contact layer, a directional capillary driving force is generated. Under the action of capillary force, exudate is automatically guided from the wound side (high wettability side) to the outside (low wettability side), achieving unidirectional fluid guidance. The small-diameter end (on the wound side) of the conical channel ensures the responsiveness of the exudate environment around the microsphere, while the large-diameter end (on the outside) reduces the resistance to exudate flow and improves the fluid conduction efficiency, thus achieving a balance between fluid conduction efficiency and responsiveness.

[0011] In this invention, the hydrophobic air-conducting layer is preferably composed of polyvinylidene fluoride (PVDF) or hydrophobically modified polycaprolactone fibers, with a fiber diameter preferably of 500-1000 nm, a contact angle ≥120°, and a water vapor transmission rate ≥3000 g / m² / 24h. This layer prevents exudate from seeping back into the wound from the outside, while allowing O2 to enter the wound and CO2 to escape, and provides overall mechanical support (tensile strength ≥5 MPa).

[0012] In this invention, the mass ratio of pH-responsive drug-loaded microspheres to hydrophilic polymer fibers in the wound contact layer is preferably 1:10 to 1:5.

[0013] In this invention, the total thickness of the gradient wetting transition layer is preferably 50-100 μm; the thickness of the wound contact layer is preferably 80-150 μm; and the thickness of the hydrophobic air-conducting layer is preferably 100-200 μm.

[0014] The present invention also provides a method for preparing the above-mentioned targeted dressing for diabetic foot infection, comprising the following steps: S1. Preparation of pH-responsive drug-loaded microspheres: Dissolve sericin in deionized water to prepare an 8-12 wt% sericin aqueous solution. Add an antibacterial drug at a drug-to-sericin mass ratio of 1:4 to 1:6. Stir at 300-500 rpm for 30-60 minutes at room temperature to ensure uniform drug dispersion. Use a reverse emulsion method, with the above sericin aqueous solution as the aqueous phase and liquid paraffin as the oil phase. Add Span 80 emulsifier, accounting for 1-3% of the oil phase mass, with a water-to-oil phase volume ratio of 1:3 to 1:6. Stir at 8000-12000 rpm. Shear emulsification at rpm for 5-10 minutes forms a stable water / oil emulsion; the emulsion is then allowed to stand and solidify at 4-10℃ for 2-4 hours, or chemical cross-linking and solidification are performed by adding glutaraldehyde (mass ratio of glutaraldehyde to sericin is 1:20 to 1:50). Microspheres are collected by centrifugation and washed 3-5 times with deionized water to obtain drug-loaded sericin microspheres; the drug-loaded microspheres are dispersed in an aqueous tannic acid solution (tannic acid concentration 3-8 mg / mL) and ultrasonically dispersed for 10-20 minutes, followed by the addition of FeCl3 solution. The tannic acid reacts with Fe... 3+ The molar ratio of tannins to Fe is 2:1 to 4:1. The mixture is stirred at 200-400 rpm for 1.5-2.5 hours at room temperature to allow the tannins to react with Fe. 3+ Interfacial coordination polymerization occurs on the surface of the microspheres to form a tannic acid-iron ion coordination polymer shell; centrifugation, washing, and freeze-drying yield pH-responsive drug-loaded microspheres. S2. Preparation of the spinning solution for the wound contact layer: Dissolve the hydrophilic polymer and the tannic acid-metal ion coordination polymer in a solvent, and disperse the drug-loaded microspheres prepared in step S1 in the spinning solution, controlling the mass ratio of microspheres to polymer to be 1:10 to 1:5. S3. Prepare gradient wetting transition layer spinning solution: Prepare polyacrylonitrile spinning solution, silk fibroin spinning solution and polycaprolactone spinning solution respectively, with the wettability of the three spinning solutions decreasing in that order. S4. Prepare the hydrophobic air-conducting layer spinning solution: Dissolve polyvinylidene fluoride or hydrophobically modified polycaprolactone in a solvent. S5. Multilayer Sequential Electrospinning and Intermediate Layer Pore Creation: A continuous liquid-changing electrospinning process is employed, sequentially electrospinning the wound contact layer and the gradient wetting transition layer to form a two-layer fiber membrane composite structure. After the gradient wetting transition layer is spun, spinning is paused, and the composite membrane is transferred to a laser processing platform. A UV solid-state laser is used to process conical channels on the gradient wetting transition layer. The laser wavelength is 355 nm, the laser power is 1-5 W, the scanning speed is 100-500 mm / s, the pulse frequency is 10-50 kHz, and the pulse width is 10-50 ns. The laser focus is controlled to be in the middle of the thickness direction of the gradient wetting transition layer using a galvanometer scanning system, allowing the conical channels to penetrate the gradient wetting transition layer. Subsequently, the composite membrane is refixed to the receiving device, and electrospinning of the hydrophobic gas-conducting layer continues, forming a three-layer fiber membrane composite structure. S6. Crosslinking stabilization treatment: The composite film obtained by spinning is subjected to steam crosslinking or ultraviolet curing treatment; S7. Post-processing: Cutting, sterilizing, and packaging to obtain the targeted dressing for diabetic foot infection.

[0015] Preferably, in step S5, the multilayer sequential electrospinning adopts a continuous liquid-changing electrospinning process, with a spinning voltage of 10-20 kV, a receiving distance of 10-20 cm, and a spinning solution flow rate of 0.5-2 mL / h.

[0016] Preferably, in step S5, the ultraviolet solid-state laser has a laser wavelength of 355 nm, a laser power of 1-5 W, a scanning speed of 100-500 mm / s, a pulse frequency of 10-50 kHz, and a pulse width of 10-50 ns. The laser focus is controlled by a galvanometer scanning system to be located in the middle of the thickness direction of the gradient wetting transition layer, so that the conical channel penetrates the gradient wetting transition layer without damaging the wound contact layer and the hydrophobic air-conducting layer. The diameter of the conical channel is 3-5 μm on the side near the wound contact layer and 15-20 μm on the side near the hydrophobic air-conducting layer. The density of the conical pores is 100-500 pores / cm².

[0017] The beneficial effects of this invention are as follows: The core innovation of this invention lies in its integrated architecture of a three-layer Janus gradient fiber membrane and embedded pH-responsive microspheres. This architecture eliminates the need for a hydrophilic fiber network in the wound contact layer to absorb exudate and provide a moist healing environment, while simultaneously using the gaps between the fibers to physically anchor the microspheres, preventing medication from being washed away by exudate. The hydrophilic fiber network and microsphere anchoring function coexist in the same layer through a "fiber-microsphere" biphase structure.

[0018] Furthermore, the gradient wetting transition layer adopts a conical pore structure. The small pore diameter end (wound side) ensures the response sensitivity of the exudate environment around the microspheres, while the large pore diameter end (outer side) ensures the fluid conduction efficiency. The fluid conduction and response are balanced through the pore size gradient rather than uniform large pores.

[0019] Furthermore, the one-way moisture-guiding mechanism continuously draws drug-containing exudate away from the wound, but the microspheres are anchored by the fiber network and are not carried away. The shell disintegration and drug release are only triggered in the infected microenvironment (pH<6.0 or high enzyme activity), while the shell stably seals the drug in normal tissue (pH≈7.4), achieving temporal decoupling of "exudate is guided away but the drug remains". Attached Figure Description

[0020] Figure 1 This is a graph comparing the cumulative drug release rates of the embodiments and comparative examples in PBS at pH 7.4 and PBS at pH 5.5.

[0021] Figure 2 This is a schematic diagram of the three-layer structure of the targeted dressing for diabetic foot infection of the present invention. Detailed Implementation

[0022] The present invention will be further described below with reference to specific embodiments, but the invention is not limited to these embodiments. Those skilled in the art should recognize that the present invention covers all possible alternatives, improvements, and equivalents included within the scope of the claims.

[0023] To make the objectives, technical solutions, and advantages of the present invention clearer, the present invention will be further described in detail below with reference to specific embodiments. Example 1

[0024] This embodiment provides a targeted dressing for diabetic foot infection, the structure of which is as follows: Figure 2 As shown, it includes a wound contact layer 1, a gradient wetting transition layer 2, and a hydrophobic air-conducting layer 3, which are stacked in sequence.

[0025] Wound contact layer 1: Composed of polycaprolactone (PCL) and tannic acid-iron ions (TA-Fe) 3+ Co-spun fibers of the coordination polymer were used as the matrix, with a fiber diameter of 300 nm and a porosity of 80%. pH-responsive drug-loaded microspheres were dispersed in the interstices of the fibers, with a mass ratio of microspheres to polymer fibers of 1:8.

[0026] The pH-responsive drug-loaded microspheres comprise: a core of porous sericin microspheres loaded with mupirocin (particle size 2-3 μm, drug loading 15 wt%), and tannic acid-iron ions (TA-Fe) coated on the surface of the core. 3+ Coordination polymer shell. In the infection microenvironment (pH < 6.0), TA-Fe... 3+Coordination bonds break, the shell disintegrates, and the drug is released; at normal pH (7.4), the shell is stable and the drug is blocked.

[0027] In this embodiment, the pH-responsive drug-loaded microspheres have a particle size of 2-3 μm, the hydrophilic polymer fibers have a diameter of 300 nm, and the fiber gaps are measured to be 150-250 nm using scanning electron microscopy. The microsphere particle size is larger than the fiber gaps, allowing the microspheres to be physically locked within the fiber network and prevent them from migrating with the exudate. Simultaneously, the microsphere particle size is smaller than the pore size (4 μm) of the conical channel on the side near the wound contact layer, ensuring that the microspheres do not clog the small-diameter end of the conical channel under exudate flushing, thus ensuring the synergy between response sensitivity and fluid conduction efficiency.

[0028] The gradient wetting transition layer 2 comprises a first transition layer (polyacrylonitrile (PAN) fiber, contact angle 40°, thickness 20 μm), a second transition layer (silk fibroin (SF) fiber, contact angle 60°, thickness 20 μm), and a third transition layer (polycaprolactone (PCL) fiber, contact angle 80°, thickness 20 μm), stacked sequentially along the direction away from the wound contact layer. The total thickness is 60 μm. The wettability of the gradient wetting transition layer 2 decreases either continuously or in a stepwise manner; this embodiment uses a stepwise decrease, meaning that the first, second, and third transition layers each have discrete contact angle ranges.

[0029] Conical channels 4 are formed through the gradient wetting transition layer 2. The diameter of the conical channels is 4 μm on the side near the wound contact layer and 18 μm on the side near the hydrophobic air-conducting layer, with a density of 300 conical channels / cm². Based on the wettability gradient driving principle, exudate is guided unidirectionally from the wound side to the outside along the direction of the arrow.

[0030] In this embodiment, the wettability gradient and pore size gradient of the gradient wetting transition layer are coupled and synergistic: the first transition layer (polyacrylonitrile fiber, contact angle 40°) is in direct contact with the wound contact layer (contact angle measured to be approximately 20°), and the contact angle difference between the two is 20°, forming a first-level capillary driving force; the small-diameter end (4 μm) of the conical channel is located near the interface between the first transition layer and the wound contact layer, and the high capillary pressure at the small-diameter end draws exudate from the wound contact layer to the first transition layer; the large-diameter end (18 μm) of the conical channel is located near the interface between the third transition layer (polycaprolactone fiber, contact angle 80°) and the hydrophobic air-conducting layer, and the low flow resistance at the large-diameter end rapidly discharges exudate to the hydrophobic air-conducting layer; the wettability gradient and the pore size gradient are in the same direction, together constituting a synergistic driving force for unidirectional fluid guidance, rather than a simple superposition of the wettability gradient or the pore size gradient acting alone.

[0031] Hydrophobic gas-conducting layer 3: Employed from polyvinylidene fluoride (PVDF) fibers with a diameter of 800 nm, a contact angle of 125°, a thickness of 150 μm, a water vapor permeability of 3500 g / m² / 24h, and a tensile strength of 8 MPa. In this embodiment, the high hydrophobicity (125° contact angle) of the hydrophobic gas-conducting layer creates a negative pressure locking effect at the large-diameter end of the conical channel: when exudate is guided to the interface of the hydrophobic gas-conducting layer, the hydrophobic surface prevents the exudate from back-permeating back into the gradient wetting transition layer, thereby forming a stable unidirectional pressure difference from the wound side to the outside within the conical channel; this negative pressure locking effect, in synergy with the physical interlocking of the microspheres in the fiber gaps, keeps the microspheres anchored in unidirectional permeation, releasing the drug only when the infection microenvironment triggers shell disintegration, achieving a triple temporal decoupling of "unidirectional exudate flow - microsphere mechanical anchoring - drug environmental response".

[0032] Preparation method: S1. Preparation of pH-responsive drug-loaded microspheres: Sericin was dissolved in deionized water to prepare a 10 wt% sericin aqueous solution. Mupirocin was added, with a drug-to-sericin mass ratio of 1:5. The mixture was stirred at 400 rpm for 45 minutes at room temperature to ensure uniform drug dispersion. Using a reverse emulsion method, the above sericin aqueous solution was used as the aqueous phase, and liquid paraffin as the oil phase. Span 80 emulsifier was added, accounting for 2% of the oil phase mass, with a water-to-oil volume ratio of 1:5. The mixture was sheared and emulsified at 10,000 rpm for 8 minutes to form a stable water / oil emulsion. The emulsion was allowed to stand at 6°C for 3 hours to solidify. The microspheres were collected by centrifugation and washed four times with deionized water to obtain drug-loaded sericin microspheres. Drug-loaded microspheres were dispersed in an aqueous solution of tannic acid (5 mg / mL) and ultrasonically dispersed for 15 minutes. Then, FeCl3 solution was added, with a molar ratio of tannic acid to Fe³⁺ of 3:1. The mixture was stirred at 300 rpm for 2 hours at room temperature to allow the tannic acid and Fe³⁺ to react. 3+ Interfacial coordination polymerization occurs on the surface of the microspheres, forming a tannic acid-iron ion coordination polymer shell. After centrifugation, washing, and freeze-drying, pH-responsive drug-loaded microspheres are obtained.

[0033] S2. Preparation of the wound contact layer spinning solution: Mix PCL (molecular weight 80,000) with TA-Fe 3+ The coordination polymer (PCL:TA = 10:1 by mass) was dissolved in hexafluoroisopropanol to prepare a 10 wt% spinning solution. The drug-loaded microspheres prepared in step S1 were ultrasonically dispersed in the spinning solution, with a microsphere to polymer mass ratio of 1:8.

[0034] S3. Prepare gradient wetting transition layer spinning solutions: Prepare three spinning solutions respectively: PAN (10 wt% in DMF), SF (12 wt% in hexafluoroisopropanol), and PCL (10 wt% in hexafluoroisopropanol).

[0035] S4. Preparation of hydrophobic and air-conducting spinning solution: Dissolve PVDF (molecular weight 300,000) in a DMF / acetone (volume ratio 7:3) mixed solvent to prepare a 12 wt% spinning solution.

[0036] S5. Multilayer Sequential Electrospinning and Intermediate Layer Pore Creation: A continuous liquid-changing electrospinning device was used, with a spinning voltage of 15kV, a receiving distance of 15 cm, and a flow rate of 1 mL / h. The following layers were spun sequentially: wound contact layer (2 hours) → first transition layer (0.5 hours) → second transition layer (0.5 hours) → third transition layer (0.5 hours), forming a two-layer composite structure of the wound contact layer and the gradient wetting transition layer. Spinning was paused, and the composite membrane was transferred to a laser processing platform. An ultraviolet solid-state laser (wavelength 355 nm, power 3 W, scanning speed 300 mm / s, pulse frequency 20 kHz, pulse width 20 ns) was used. The laser focus was controlled to be at the center of the gradient wetting transition layer thickness using a galvanometer scanning system to process conical channels. The diameter of the conical channels near the wound contact layer was 4 μm, and the diameter near the hydrophobic gas-conducting layer was 18 μm, with a conical channel density of 300 channels / cm². The composite membrane was then reattached to the receiving device, and the spinning of the hydrophobic air-conducting layer continued (1.5 hours) to form a three-layer composite fiber membrane. The ultraviolet solid-state laser emitted from the side of the gradient wetting transition layer away from the wound contact layer towards the wound contact layer, with the laser focus positioned at the middle of the thickness direction of the gradient wetting transition layer. Because the polyacrylonitrile, silk fibroin, and polycaprolactone in the gradient wetting transition layer significantly absorb 355 nm ultraviolet laser light, while the tannic acid-metal ion coordination polymer in the wound contact layer and the polyvinylidene fluoride in the hydrophobic air-conducting layer have low absorption rates for this wavelength, and because the laser energy is concentrated at the focal point and rapidly attenuates outside the focal point, an ablation channel is formed only within the gradient wetting transition layer without damaging the wound contact layer and the hydrophobic air-conducting layer. Simultaneously, utilizing the energy gradient distribution within the Rayleigh length range before and after the laser focal point, a tapered channel with a gradually changing aperture from the focal point to both sides is naturally formed. S6. Crosslinking stabilization treatment: The composite membrane is placed in glutaraldehyde vapor for 2 hours to crosslink, which enhances fiber stability and interlayer bonding.

[0037] S7. Post-processing: Cut to 10 cm × 10 cm specifications, sterilize with ethylene oxide, and seal in packaging.

[0038] Performance testing: (1) Drug release test: The dressing of Example 1 was placed in PBS buffer at pH 7.4 and pH 5.5 respectively and shaken at 37°C. The results showed that the drug release was 8.2% after 24 h at pH 7.4 and 75.6% after 12 h at pH 5.5.

[0039] (2) Unidirectional fluid infiltration test: A modified vertical fluid infiltration test device was used to measure the permeation volume in the forward (wound surface → outside) and reverse (outside → wound surface) directions. The unidirectional fluid infiltration coefficient (forward / reverse ratio) of Example 1 was 6.8.

[0040] (3) Antibacterial test: The diameter of the inhibition zone against Staphylococcus aureus (ATCC 25923) was 28 mm, and the diameter of the inhibition zone against Pseudomonas aeruginosa (ATCC 27853) was 24 mm.

[0041] (4) Unidirectional fluid infiltration test method: A modified vertical fluid infiltration test device was used, consisting of two independent liquid tanks, upper and lower, sealed and isolated by a horizontally placed dressing sample. The lower liquid tank contained pH 7.4 PBS buffer, with the liquid level kept constant at 2 cm; the upper liquid tank was initially empty. The mass of fluid infiltrating in both the forward (bottom → top, simulated wound → outside) and reverse (top → bottom) directions was measured over 24 h. The unidirectional fluid infiltration coefficient was defined as the ratio of forward infiltration to reverse infiltration.

[0042] (5) Effective antibacterial time test method: Apply the dressing sample to the inoculated Staphylococcus aureus (1×10⁻⁶) sample. 6 Inoculate the plates with CFU / mL on agar plates and incubate at 37°C. Replace the plates every 24 hours and observe and record the diameter of the inhibition zone. An inhibition zone diameter ≥10 mm is considered to have antibacterial activity, and record the number of days the antibacterial activity lasts.

[0043] (6) Microsphere anchoring force test: The dressing of Example 1 was fixed in a vertical percolation device and flushed from the wound contact layer to the hydrophobic air-conducting layer with pH 7.4 PBS buffer at a flow rate of 0.5 mL / min for 24 h. The effluent was collected and centrifuged to detect the microsphere content. The results showed that the cumulative microsphere loss rate in 24 h was <0.5%, proving that the physical interlocking of microspheres in the fiber gaps and the negative pressure locking effect of the hydrophobic air-conducting layer can effectively prevent microspheres from being washed away by the exudate.

[0044] (7) Photothermal-assisted antibacterial test: The dressing of Example 1 was placed under irradiation with an 808 nm near-infrared laser (power density 0.8 W / cm², irradiation for 5 min), and the surface temperature change was recorded using an infrared thermal imager. The results showed that the surface temperature of the dressing increased from 37℃ to 50℃ (ΔT=13℃), and the diameter of the inhibition zone against Staphylococcus aureus increased to 34 mm under light conditions (28 mm under no light conditions), proving that the tannic acid-iron ion complex endows the fiber with photothermal conversion ability, which can achieve photothermal-drug synergistic antibacterial effect. Example 2

[0045] The difference from Example 1 is that the mass ratio of microspheres to polymer fibers in the wound contact layer is 1:5; the total thickness of the gradient wetting transition layer is 80 μm (the thicknesses of each layer are 25 μm, 25 μm, and 30 μm, respectively); and the hydrophobic air-conducting layer uses hydrophobically modified PCL (grafted perfluoroalkyl chain) instead of PVDF.

[0046] The rest is the same as in Example 1.

[0047] Performance testing: The drug release rate was 6.5% after 24 hours at pH 7.4; the drug release rate was 82.3% after 12 hours at pH 5.5; and the unidirectional liquid-carrying coefficient was 7.2. In this embodiment, the three-level size anchoring relationship between the pH-responsive drug-loaded microspheres, hydrophilic polymer fibers, and conical channels, the coupling synergistic relationship between the wettability gradient and the pore size gradient, and the negative pressure locking time-sequence decoupling mechanism of the hydrophobic gas-conducting layer are all the same as in Example 1, and will not be repeated here. Example 3

[0048] The difference from Example 1 is that the metal ions in the pH-responsive drug-loaded microspheres are Cu. 2+ Fe replacement 3+ The antibacterial drug uses silver nanoparticles (particle size 20 nm, drug loading 10 wt%) instead of mupirocin; the density of the conical pores is 500 / cm².

[0049] The rest is the same as in Example 1.

[0050] Performance testing: Drug release was 9.1% after 24 hours at pH 7.4; 71.4% after 12 hours at pH 5.5; unidirectional fluid conduction coefficient was 8.5; and the inhibition zone diameter against Staphylococcus aureus was 31 mm. In this embodiment, the three-level dimensional anchoring relationship between the pH-responsive drug-loaded microspheres, hydrophilic polymer fibers, and conical channels, the coupling synergistic relationship between the wettability gradient and the pore size gradient, and the negative pressure locking time-sequence decoupling mechanism of the hydrophobic air-conducting layer are all the same as in Example 1, and will not be repeated here.

[0051] Comparative Example 1 (without microsphere anchoring structure) The same Janus gradient fiber membrane structure as in Example 1 was used, but instead of embedding pH-responsive drug-loaded microspheres in the wound contact layer, mupirocin was dissolved in a small amount of DMSO and dispersed in the spinning solution. The rest was the same as in Example 1.

[0052] Performance testing: Drug release was 65.3% after 24 hours at pH 7.4; 88.7% after 12 hours at pH 5.5; the unidirectional fluidization coefficient was 6.5. Due to the lack of a protective shell and microsphere anchoring, the drug was rapidly washed away by exudate, failing to achieve targeted release for infection.

[0053] Comparative Example 2 (No Gradient Wetting Transition Layer) The same wound contact layer and hydrophobic air-conducting layer as in Example 1 were used, but the intermediate layer was a uniform PCL fiber layer (without wettability gradient and without conical channels). The rest was the same as in Example 1.

[0054] Performance test: The one-way fluid guiding coefficient is 1.2 (almost no one-way fluid guiding effect), the wound surface contact area is severely affected by fluid, and the wound surface humidity reaches more than 95% after 24 hours, resulting in maceration.

[0055] Comparative Example 3 (Traditional Foam Dressing) Commercially available polyurethane foam dressing (containing silver ions) with a thickness of 3 mm was used.

[0056] Performance testing: No pH-responsive drug release capability; 24-hour drug release was 45% (no difference at pH 7.4 and pH 5.5); no unidirectional fluid guiding capability; exudate diffuses bidirectionally; drug change cycle is 3 days.

[0057] Comparative analysis: As shown in Tables 1 and 2, Examples 1-3 all exhibited low drug release (<10%) at pH 7.4, but rapid release (>70%) at pH 5.5, showing obvious infection targeting; Comparative Example 1, lacking shell protection, released rapidly at both pH levels, showing no targeting; Comparative Example 3 showed continuous release, with no environmental responsiveness.

[0058] As shown in Table 3, the unidirectional fluid guiding coefficients of Examples 1-3 are all greater than 5, which is significantly better than Comparative Example 2 (1.2) and Comparative Example 3 (without unidirectional fluid guiding), proving that the tapered channel gradient structure is the key to achieving unidirectional fluid guiding.

[0059] The cumulative release rate (%) of pH 7.4 PBS was compared and analyzed in Table 1.

[0060] Data characteristics: Example 1: Endpoint release rate 8.2%, stable shell, drug occlusion. Comparative Example 1: Endpoint release rate 65.3%, no shell protection, severe burst release. Comparative Example 2: Endpoint release rate 32.5%, shell present but wound maceration accelerates diffusion. Comparative Example 3: Endpoint release rate 45.0%, conventional foam release, no pH response.

[0061] The cumulative release rate (%) of PBS at pH 5.5 is compared and analyzed in Table 2 below:

[0062] Data characteristics: Example 1 showed a final release rate of 75.6%, with shell disintegration and pulsatile release. Comparative Example 1 showed a final release rate of 88.7%, with no shell and rapid loss. Comparative Example 2 showed a final release rate of 79.3%, with shell disintegration and impregnation, and release slightly faster than Example 1. Comparative Example 3 showed a final release rate of 42.5%, with continuous release and no environmental response (almost identical to the data at pH 7.4). At pH 5.5, the drug was essentially completely released within 12 hours, therefore 24-hour data were not measured.

[0063] Table 3. Comparison of unidirectional fluid guide coefficient and antibacterial performance:

[0064] Note: "—" indicates that the performance was not tested or is not available. Comparative Example 2, lacking unidirectional fluid guiding ability, showed maceration of the wound side with a moisture content exceeding 95% after 24 hours. Comparative Example 3, a commercially available traditional foam dressing, also lacked unidirectional fluid guiding ability, resulting in bidirectional diffusion of exudate.

[0065] With the approval of the medical institution's ethics committee and the informed consent of the patients, the dressing of Example 1 was used on patients with diabetic foot ulcers (Wagner grade 2-3) for a 7-day treatment period. Results showed that wound exudate was effectively diverted to the outer absorbent pad, maintaining a relative humidity of 85%-90% on the wound contact side (neither dry nor soaked); during the infection phase (pH < 6.5), the drug was automatically released, resulting in a two-order-of-magnitude decrease in bacterial count; during the non-infection phase, the drug was almost not released, preventing the development of drug resistance. A single application provided effective antibacterial protection for more than 7 days, more than double the duration of traditional silver-containing dressings (3 days).

[0066] Compared with existing technologies, this invention has the following advantages: Through a biphasic structure of hydrophilic fiber network and embedded microspheres, "hydrophilic moisture conduction" and "microsphere drug anchoring" coexist in the same wound contact layer, preventing drug loss due to exudate washout. The conical pore structure of the gradient wetting transition layer utilizes the wettability gradient to drive unidirectional exudate outflow, while the small pore size ensures microsphere responsiveness, achieving a balance between drainage efficiency and responsiveness. The pH-responsive shell design achieves temporal decoupling of "exudate removal but drug retention," triggering drug release only in the infected microenvironment and blocking the drug under normal conditions, thus preventing drug resistance.

[0067] It also extends the dressing change cycle: the effective antibacterial time of a single dressing is ≥7 days, which is 2-3 times longer than traditional silver-containing dressings, reducing the frequency of dressing changes and the risk of secondary infection. The wound side is maintained at 85%-95% relative humidity, while excess exudate is drained in time, meeting the dual standards of "not drying and not soaking".

[0068] The above description is only a preferred embodiment of the present invention, but the scope of protection of the present invention is not limited thereto. Any equivalent substitutions or modifications made by those skilled in the art within the scope of the technology disclosed in the present invention, based on the technical solution and inventive concept of the present invention, should be covered within the scope of protection of the present invention.

Claims

1. A targeted dressing for diabetic foot infection, characterized in that, It includes a wound contact layer, a gradient wetting transition layer, and a hydrophobic air-conducting layer, which are stacked in sequence; The wound contact layer is a hydrophilic microsphere embedded in a fiber layer, consisting of a fiber network formed by interwoven hydrophilic polymer fibers and pH-responsive drug-loaded microspheres with a particle size larger than the pores of the fiber network and physically locked therein; the pH-responsive drug-loaded microspheres include: a core of porous sericin microspheres loaded with antibacterial drugs, and a tannic acid-metal ion coordination polymer shell coating the surface of the core. The gradient wetting transition layer is a multi-layer fiber structure in which the fiber wettability decreases from the wound contact layer to the hydrophobic air-conducting layer, and a conical channel is provided through the gradient wetting transition layer. The diameter of the conical channel gradually increases from the side closer to the wound contact layer to the side closer to the hydrophobic air-conducting layer. The hydrophobic air-conducting layer is a hydrophobic polymer fiber layer.

2. The targeted dressing for diabetic foot infection according to claim 1, characterized in that, The hydrophilic polymer fiber of the wound contact layer is a co-spun fiber of polycaprolactone and tannic acid-metal ion coordination polymer, with a fiber diameter of 200-500 nm and a porosity of 70%-85%.

3. The targeted dressing for diabetic foot infection according to claim 1, characterized in that, In the pH-responsive drug-loaded microspheres, the metal ions are selected from Fe. 3+ Cu 2+ At least one of the following: the antibacterial drug is selected from at least one of mupirocin and silver nanoparticles; the microspheres have a particle size of 1-5 μm.

4. The targeted dressing for diabetic foot infection according to claim 1, characterized in that, The gradient wetting transition layer comprises layers stacked sequentially along the direction away from the wound contact layer: The first transition layer is composed of polyacrylonitrile fibers with a contact angle of 30°-50°. The second transition layer is composed of silk fibroin fibers with a contact angle of 50°-70°. The third transition layer is composed of polycaprolactone fibers with a contact angle of 70°-90°. Furthermore, the diameter of the conical channel is 3-5 μm on the side near the wound contact layer and 15-20 μm on the side near the hydrophobic air-conducting layer, with a conical pore density of 100-500 pores / cm².

5. The targeted dressing for diabetic foot infection according to claim 1, characterized in that, The hydrophobic gas-conducting layer is composed of polyvinylidene fluoride or hydrophobically modified polycaprolactone fibers with a fiber diameter of 500-1000 nm, a contact angle ≥120°, and a water vapor transmission rate ≥3000 g / m² / 24h.

6. The targeted dressing for diabetic foot infection according to claim 1, characterized in that, The mass ratio of pH-responsive drug-loaded microspheres to hydrophilic polymer fibers in the wound contact layer is 1:10 to 1:

5.

7. The targeted dressing for diabetic foot infection according to claim 1, characterized in that, The total thickness of the gradient wetting transition layer is 50-100 μm; the thickness of the wound contact layer is 80-150 μm; and the thickness of the hydrophobic air-conducting layer is 100-200 μm.

8. A method for preparing a targeted dressing for diabetic foot infection as described in any one of claims 1-7, characterized in that, Includes the following steps: S1. Preparation of pH-responsive drug-loaded microspheres: Dissolve sericin in deionized water to prepare an 8-12 wt% sericin aqueous solution, add antibacterial drugs, with a drug-to-sericin mass ratio of 1:4 to 1:6, and stir at 300-500 rpm for 30-60 minutes at room temperature to ensure uniform drug dispersion; use the reverse emulsion method, with the above sericin aqueous solution as the aqueous phase and liquid paraffin as the oil phase, add emulsifier, and shear emulsify at 8000-12000 rpm for 5-10 minutes to form a stable water / oil emulsion; The emulsion was allowed to stand and solidify at 4-10℃, or chemically cross-linked and solidified by adding glutaraldehyde. Microspheres were collected by centrifugation and washed 3-5 times with deionized water to obtain drug-loaded sericin microspheres. The drug-loaded microspheres were dispersed in an aqueous tannic acid solution (3-8 mg / mL) and ultrasonically dispersed for 10-20 minutes. Subsequently, FeCl3 solution was added, and the tannic acid reacted with Fe... 3+ The molar ratio of tannic acid to Fe was 2:1 to 4:1, and the mixture was stirred at room temperature to allow the tannic acid to react with Fe. 3+ Interfacial coordination polymerization occurs on the surface of the microspheres to form a tannic acid-iron ion coordination polymer shell; centrifugation, washing, and freeze-drying yield pH-responsive drug-loaded microspheres. S2. Preparation of the spinning solution for the wound contact layer: Dissolve the hydrophilic polymer and the tannic acid-metal ion coordination polymer in a solvent, and disperse the drug-loaded microspheres prepared in step S1 in the spinning solution, controlling the mass ratio of microspheres to polymer to be 1:10 to 1:

5. S3. Prepare gradient wetting transition layer spinning solution: Prepare polyacrylonitrile spinning solution, silk fibroin spinning solution and polycaprolactone spinning solution respectively, with the wettability of the three spinning solutions decreasing in that order. S4. Prepare the hydrophobic air-conducting layer spinning solution: Dissolve polyvinylidene fluoride or hydrophobically modified polycaprolactone in a solvent. S5. Multilayer Sequential Electrospinning and Intermediate Layer Pore Creation: A continuous liquid-changing electrospinning process is employed, sequentially electrospinning the wound contact layer and the gradient wetting transition layer to form a two-layer fiber membrane composite structure. After the gradient wetting transition layer is spun, spinning is paused, and the composite membrane is transferred to a laser processing platform. A UV solid-state laser is used to process conical channels on the gradient wetting transition layer. The laser focus is controlled to be in the middle of the thickness direction of the gradient wetting transition layer by a galvanometer scanning system, allowing the conical channels to penetrate the gradient wetting transition layer. Subsequently, the composite membrane is refixed to the receiving device, and electrospinning of the hydrophobic gas-conducting layer continues, forming a three-layer fiber membrane composite structure. S6. Crosslinking stabilization treatment: The composite film obtained by spinning is subjected to steam crosslinking or ultraviolet curing treatment; S7. Post-processing: Cutting, sterilizing, and packaging to obtain the targeted dressing for diabetic foot infection.

9. The preparation method according to claim 8, characterized in that, In step S5, the multilayer sequential electrospinning adopts a continuous liquid-changing electrospinning process, with a spinning voltage of 10-20 kV, a receiving distance of 10-20 cm, and a spinning solution flow rate of 0.5-2 mL / h. The ultraviolet solid-state laser emits light from the side of the gradient wetting transition layer away from the wound contact layer toward the wound contact layer. The laser focus is located in the middle of the thickness direction of the gradient wetting transition layer. The laser wavelength of the ultraviolet solid-state laser is 355 nm, the laser power is 1-5 W, the scanning speed is 100-500 mm / s, the pulse frequency is 10-50 kHz, and the pulse width is 10-50 ns. The laser focus is controlled to be in the middle of the thickness direction of the gradient wetting transition layer by a galvanometer scanning system, so that the conical channel penetrates the gradient wetting transition layer without damaging the wound contact layer and the hydrophobic gas-conducting layer.

10. The preparation method according to claim 9, characterized in that, The emulsifier in S1 is Span 80, which accounts for 1-3% of the oil phase mass, and the volume ratio of the water phase to the oil phase is 1:3 to 1:

6. When static curing is used, the curing time is 2-4 hours. When glutaraldehyde chemical crosslinking curing is used, the mass ratio of glutaraldehyde to sericin is 1:20 to 1:50.

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