Anisotropic drug-loading core-shell fiber scaffold and preparation method thereof

Abstract

The invention belongs to the technical field of biomedicine, and discloses an anisotropic drug-loading core-shell fiber scaffold and a preparation method thereof.The anisotropic drug-loading core-shell fiber scaffold sequentially comprises a core layer, a first shell layer and a second shell layer; the core layer comprises a porous core layer matrix and a hydrophilic drug loaded on the porous core layer matrix; the first shell layer comprises a porous shell layer matrix and a hydrophobic drug loaded on the porous shell layer matrix; the second shell layer is obtained by crosslinking and curing modified gelatin through genipin. Through ordered configuration design of the core layer, the first shell layer and the second shell layer, partitioned loading and independent controlled release of a hydrophilic drug and a hydrophobic drug and collaborative optimization of scaffold biological activity are realized, and the technical problems of poor compatibility of multiple drugs, uncontrollable release, insufficient biological activity and the like in an existing drug-loaded fiber scaffold are solved from the structural design level; and the stent has excellent anisotropy, and a functional stent with targeting drug release and biocompatibility is provided for complex treatment scenes.

Description

Technical Field

[0001] This invention belongs to the field of biomedical technology, and specifically relates to an anisotropic drug-loaded core-shell fiber scaffold and its preparation method. Background Technology

[0002] Drug-loaded fiber scaffolds prepared by electrospinning are widely used in biomedical fields such as tissue repair, anti-inflammation, and anticoagulation due to their ability to mimic the natural extracellular matrix and their ability to release drugs sustainably, making them an important research direction for meeting complex treatment needs. However, in current technologies, traditional uniaxial electrospun fibers can only achieve single-drug loading. If multiple drugs with different physicochemical properties are loaded simultaneously, compatibility problems such as aggregation and mutual inhibition of activity can easily occur due to direct contact between drugs. Furthermore, it is impossible to achieve independent release regulation of different drugs, making it difficult to meet the requirements of sequential and targeted release of multiple drugs in complex treatment scenarios.

[0003] To address the challenge of multi-drug loading, coaxial electrospinning technology, while capable of fabricating core-shell structured fibers to achieve physical drug isolation, often suffers from delamination and drug leakage due to insufficient polymer interfacial compatibility, affecting structural stability and drug release efficacy. Meanwhile, while synthetic polymers commonly used in tissue engineering, such as polylactic acid and polycaprolactone, possess good biocompatibility, biodegradability, and mechanical properties, they lack bioactive sites, resulting in poor cell adhesion and difficulty in supporting the cell proliferation microenvironment required for tissue repair.

[0004] Although some studies have attempted to modify natural polymers such as gelatin to enhance bioactivity, natural gelatin has poor stability, and traditional cross-linking methods either pose a risk of toxic residues or have poor cross-linking effects, failing to balance bioactivity retention and structural stability. Furthermore, it is difficult to effectively control the initial drug release rate, which can easily lead to drug burst release, further limiting the clinical application value of drug-loaded stents. Therefore, the development of drug-loaded stents with multi-drug zone loading, independent controlled release, interfacial stability, and high bioactivity has become an urgent need in the field. Summary of the Invention

[0005] The present invention aims to improve at least one technical problem in the prior art.

[0006] The first aspect of the present invention provides an anisotropic drug-loaded core-shell fiber scaffold, comprising a core layer, a first shell layer, and a second shell layer in sequence; The core layer includes a porous core layer matrix and a hydrophilic drug loaded on the porous core layer matrix, wherein the porous core layer matrix is ​​formed of polycaprolactone; The first shell includes a porous shell matrix and a hydrophobic drug loaded on the porous shell matrix, wherein the porous shell matrix is ​​formed of polylactic acid; The second shell layer is obtained by cross-linking and curing modified gelatin with genipin, and the modified gelatin is obtained by methacrylating gelatin.

[0007] The porous core layer matrix of the anisotropic drug-loaded core-shell fiber scaffold provided in this application is formed of polycaprolactone. Polycaprolactone is an excellent biocompatible polymer with a suitable degradation rate matching the human tissue repair cycle, good mechanical strength, and excellent molecular chain structure characteristics. The ester bonds contained in the polycaprolactone molecular chain are easily hydrolyzed and degraded, and the crystallinity is moderate, which can form a microenvironment conducive to the dispersion of hydrophilic drugs and avoid drug aggregation. At the same time, polycaprolactone has good compatibility with hydrophilic drugs (such as heparin sodium) and can achieve stable loading of hydrophilic drugs through physical adsorption and intermolecular forces, laying the foundation for long-acting and controllable drug release. The first shell layer surrounds the core layer, forming the intermediate sheath layer of the fibrous scaffold. The matrix of the first shell layer (porous shell matrix) is made of polylactic acid (PLA), which possesses excellent biocompatibility and biodegradability. PLA also exhibits good polymer compatibility with polycaprolactone (PVC) in the core layer. Due to their similar molecular chain structures, molecular chain entanglement can be achieved through co-solvent spinning, forming a stable and continuous core-shell interface. This avoids the delamination and drug leakage problems caused by interfacial incompatibility in traditional core-shell structures. The hydrophobic drug (such as curcumin) loaded in the first shell layer is compatible with the hydrophobic properties of PLA. The hydrophobic drug can be uniformly dispersed in the PLA matrix, achieving physical isolation from the hydrophilic drug in the core layer. This partitioned loading design fundamentally solves the problems of aggregation, inactivation, or mutual inhibition caused by direct contact between hydrophilic and hydrophobic drugs, ensuring that the two drugs do not interfere with each other during storage and release. The second shell, as the outermost structure of the anisotropic drug-loaded core-shell fiber scaffold of this application, is obtained by cross-linking and curing methacrylated modified gelatin with genipin. Gelatin is a natural biological macromolecule with excellent biocompatibility and cell adhesion, but natural gelatin has poor stability and is easily degraded. This application modifies the structure of gelatin by methacrylation modification, introducing methacryloyl groups into the gelatin molecular chain, which not only retains the original biological active sites of gelatin, but also significantly improves the chemical stability and reactivity of gelatin, providing reaction sites for subsequent cross-linking and curing. This application uses genipin as a cross-linking agent instead of a traditional chemical cross-linking agent. The core advantage of genipin lies in its being a natural plant extract with excellent biocompatibility and no cytotoxicity. The cross-linking reaction is mild yet highly efficient. Through the reaction of genipin with the active groups on the modified gelatin molecular chain, a stable cross-linked network structure is formed. This allows the second shell layer to tightly cover the surface of the first shell layer and the porous structure of the fibrous scaffold. This prevents the scaffold from rapidly swelling and detaching in the body fluid environment and also regulates the initial drug release rate through the porous structure of the cross-linked network, reducing the initial burst release effect. Simultaneously, this layer provides abundant bioactive sites, promoting cell adhesion and proliferation, and creating a favorable microenvironment for tissue repair.

[0008] The anisotropic drug-loaded core-shell fiber scaffold of this application features a three-layer core-shell structure that is not a simple stacking combination, but rather an integration of multiple technological advantages through the synergy of material selection and structural design. The polycaprolactone-polylactic acid combination of the core layer and the first shell enables partitioned loading and independent controlled release of hydrophilic and hydrophobic drugs, solving the compatibility problem of multiple drugs. The methacryloyl gelatin-genipin crosslinking design of the second shell compensates for the insufficient bioactivity of the synthetic polymer while optimizing drug release behavior. The interfacial stability of the three-layer structure, including the molecular chain entanglement of polylactic acid and polycaprolactone, and the tight bonding between the second and first shells, ensures the structural integrity of the fiber scaffold during preparation, storage, and application, providing dual protection for drug controlled release and biocompatibility, making it suitable for various complex therapeutic scenarios such as tissue repair, anti-inflammatory and anticoagulant treatments.

[0009] In some preferred embodiments, the method for preparing modified gelatin includes the following steps: Gelatin was dissolved in a carbonate buffer solution, the pH was adjusted to 8-10, and methacrylic anhydride was added under the condition of heating to 40℃-60℃. After stirring and reacting, the pH was adjusted to 6-8, dialyzed, and freeze-dried to obtain modified gelatin.

[0010] The mass-to-volume ratio of gelatin to methacrylic anhydride is 10g:(0.5-2)mL.

[0011] The concentration of gelatin in the carbonate buffer solution is 1wt%-20wt%, and the concentration of the carbonate buffer solution is 0.1M-0.5M.

[0012] The stirring speed for the reaction is 200 rpm to 1000 rpm, and the dialysis time is 2 to 4 days.

[0013] A second aspect of the present invention provides a method for preparing the above-mentioned anisotropic drug-loaded core-shell fiber scaffold, wherein the modified gelatin solution is prepared from modified gelatin and the genipin solution is prepared from genipin; the method for preparing the anisotropic drug-loaded core-shell fiber scaffold includes the following steps: Polylactic acid and a hydrophobic drug are added to the first solvent to prepare spinning solution A; Polycaprolactone and a hydrophilic drug are added to a second solvent to prepare spinning solution B; Spinning solution A and spinning solution B are respectively loaded into syringes and coaxial electrospinning is performed to obtain electrospun materials. First, the electrospun fabric is hydrophilized in a Tween 80 solution, then adsorbed into a modified gelatin solution, and finally cross-linked and cured in an aqueous solution of genipin to obtain an anisotropic drug-loaded core-shell fiber scaffold.

[0014] The first solvent comprises dichloromethane, hexafluoroisopropanol and anhydrous ethanol in a volume ratio of 0.85:0.15:(0-0.06), and the second solvent comprises dichloromethane, hexafluoroisopropanol and anhydrous ethanol in a volume ratio of 0.85:0.15:(0-0.06).

[0015] The spinning solution A has a mass fraction of 11wt%-15wt%, and the spinning solution B has a mass fraction of 11wt%-15wt%. The propulsion speed of spinning solution A is 0.5 mL / h-2 mL / h, and the propulsion speed of spinning solution B is 0.1 mL / h-0.5 mL / h.

[0016] The mass ratio of polylactic acid to hydrophobic drug is 1:(0.01-0.05).

[0017] The mass ratio of polycaprolactone to the hydrophilic drug is 1:(0.01-0.05).

[0018] The Tween 80 solution had a mass fraction of 0.05wt%-0.2wt%, and the hydrophilization treatment time was 30min-60min.

[0019] The modified gelatin solution had a mass fraction of 1wt%-5wt% and an adsorption time of 30min-120min.

[0020] The mass fraction of the genipin solution was 0.05wt%-0.2wt%, and the cross-linking curing time was 2h.

[0021] The coaxial electrospinning time is 180min-360min, and the coaxial electrospinning collection speed is 1000rpm-2000rpm.

[0022] The positive voltage of coaxial electrospinning is 10kV-15kV, the negative voltage is 0kV-3kV, the receiving distance is 10cm-20cm, the ambient temperature is controlled at 25℃, and the relative humidity is maintained at 50%.

[0023] The beneficial effects of the present invention: The anisotropic drug-loaded core-shell fiber scaffold of the present invention achieves partitioned loading, independent controlled release and synergistic optimization of scaffold bioactivity of hydrophilic and hydrophobic drugs through the ordered configuration design of the core layer, the first shell layer and the second shell layer. It solves the technical problems of poor compatibility of multiple drugs, uncontrollable release and insufficient bioactivity in existing drug-loaded fiber scaffolds from the structural design level, and provides a functional scaffold with both targeted drug release and biocompatibility for complex treatment scenarios. Attached Figure Description

[0024] To more clearly illustrate the specific embodiments of the present invention or the technical solutions in the prior art, the drawings used in the description of the specific embodiments or the prior art will be briefly introduced below. Obviously, the drawings described below are some embodiments of the present invention. For those skilled in the art, other drawings can be obtained from these drawings without creative effort.

[0025] Figure 1 This is a schematic diagram illustrating the fabrication of an anisotropic drug-loaded core-shell fiber scaffold in one embodiment. Figure 2 A scanning electron microscope image of product 1 prepared for the example; Figure 3 Scanning electron microscope image of product 2 obtained in the example; Figure 4 Scanning electron microscope image of product 3 obtained in the example; Figure 5 Scanning electron microscope image of product 4 obtained in the example; Figure 6 A scanning electron microscope image of product 5 obtained in the example; Figure 7 A scanning electron microscope image of product 6 obtained in the example; Figure 8 A scanning electron microscope image of product 7 obtained in the example; Figure 9 Transmission electron microscope image of product 1 prepared for example; Figure 10 Transmission electron microscope image of product 7 prepared for example; Figure 11 Drug release profiles of products 1 and 7 prepared for the example; Figure 12 The diagram shows the effect of product 1 and products 8-13 prepared for the example on the proliferation of human adipose-derived mesenchymal stem cells. Detailed Implementation

[0026] The present invention will be further described below with reference to specific embodiments. It should be understood that these embodiments are for illustrative purposes only and are not intended to limit the scope of the invention. Furthermore, it should be understood that after reading the description of this invention, those skilled in the art can make various alterations or modifications to the invention, and these equivalent forms also fall within the scope defined by the appended claims.

[0027] Example 1 An anisotropic drug-loaded core-shell fiber scaffold comprises, in sequence, a core layer, a first shell layer, and a second shell layer.

[0028] The core layer comprises polycaprolactone (a porous core matrix) and sodium heparin (a hydrophilic drug) loaded with polycaprolactone; the first shell layer comprises polylactic acid (a porous shell matrix) and curcumin (a hydrophobic drug) loaded with polylactic acid; the second shell layer is obtained by cross-linking and curdling modified gelatin with genipin, and the modified gelatin is obtained by methacrylating gelatin.

[0029] The preparation method of modified gelatin includes the following steps: Stirring at 50°C, 10g of gelatin was dissolved in 0.25M carbonate buffer solution (prepared by weighing 0.318g of sodium carbonate and 0.586g of sodium bicarbonate and dissolving them in 100 mL of ultrapure water) to prepare a 10wt% gelatin solution. The pH of the gelatin solution was adjusted to 9 using 5 mol / L sodium hydroxide solution and 6 mol / L hydrochloric acid solution. At 50°C, 1 mL of methacrylic anhydride (mass-volume ratio of gelatin to methacrylic anhydride is 10 g: 1 mL) was added to the gelatin solution. After stirring at 500 rpm for 1 hour, the pH was adjusted to 7.5 to end the reaction. The product of the above stirring reaction was centrifuged at 8000 rpm for 5 min to remove unreacted precipitate, then dialyzed for 3 days using a dialysis bag with a molecular weight cutoff of 3500, and freeze-dried to obtain modified gelatin.

[0030] refer to Figure 1 The method for preparing this anisotropic drug-loaded core-shell fiber scaffold includes the following steps: 1 g of polylactic acid and 0.01 g of curcumin (the mass ratio of polylactic acid to the hydrophobic drug curcumin is 1:0.01) were added to 7.5 mL of the first solvent (prepared by mixing dichloromethane, hexafluoroisopropanol and anhydrous ethanol in a volume ratio of 0.85:0.15:0.03) to prepare a spinning solution A with a mass fraction of 13 wt%. 1 g of polycaprolactone and 0.03 g of sodium heparin (the mass ratio of polycaprolactone to the hydrophilic drug sodium heparin is 1:0.03) were added to 7.5 mL of a second solvent (prepared by mixing dichloromethane, hexafluoroisopropanol and anhydrous ethanol in a volume ratio of 0.85:0.15:0.03) to prepare a spinning solution B with a mass fraction of 13 wt%. Spinning solution A and spinning solution B were respectively loaded into 10mL syringes and coaxial electrospinning was performed (ambient temperature controlled at 25℃, relative humidity maintained at 50%; the feed rate of spinning solution A was 1.5mL / h, and the feed rate of spinning solution B was 0.3mL / h; the parameters of the electrospinning device were set as follows: positive voltage 15kV, negative voltage 3kV, receiving distance 15cm, needle size 17 for spinning solution A, needle size 22 for spinning solution B, roller collection speed 2000rpm, coaxial electrospinning for 240min) to obtain electrospun material; First, the electrospun fabric was placed in a 0.05 wt% Tween 80 solution for hydrophilization treatment (shaking in the 0.05 wt% Tween 80 solution on a shaker for 45 min). After rinsing with deionized water and thoroughly cleaning with ultrasound, it was placed in a 3 wt% modified gelatin solution for modified gelatin adsorption (allowed to stand for 60 min for adsorption). Finally, it was placed in a 0.1 wt% genipin solution for cross-linking and curing for 2 h. After thorough washing with deionized water and ethanol to remove all residues, an anisotropic drug-loaded core-shell fiber scaffold was obtained.

[0031] The anisotropic drug-loaded core-shell fiber scaffold prepared in Example 1 is designated as Product 1.

[0032] Example 2 An anisotropic drug-loaded core-shell fiber scaffold differs from Example 1 in that the amounts of the first and second solvents are different: 10.5 mL / 20.5 mL of the first solvent, 9.5 mL / 29.5 mL of the first solvent, and 8.5 mL / 28.5 mL of the first solvent, respectively. This results in spinning solutions A and B having mass fractions of 9.5 wt% / 9.5 wt%, 10.5 wt% / 10.5 wt% and 11.8 wt% / 11.8 wt%, respectively.

[0033] The anisotropic drug-loaded core-shell fiber scaffolds prepared in Example 2 are referred to as Product 2, Product 3, and Product 4, respectively.

[0034] Example 3 An anisotropic drug-loaded core-shell fiber scaffold differs from Example 1 in that the volume ratio of anhydrous ethanol in the first solvent and the second solvent is different. Specifically, the first solvent (prepared from dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0) / the second solvent (prepared from dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0) / the second solvent (prepared from dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0) / the second solvent (prepared from dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0). The first solvent (prepared by mixing dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0.06) and the second solvent (prepared by mixing dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0.06) and the third solvent (prepared by mixing dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0.09) and the fourth solvent (prepared by mixing dichloromethane, hexafluoroisopropanol, and anhydrous ethanol in a volume ratio of 0.85:0.15:0.09).

[0035] The anisotropic drug-loaded core-shell fiber scaffolds prepared in Example 3 are referred to as Product 5, Product 6, and Product 7, respectively.

[0036] Example 4 An anisotropic drug-loaded core-shell fiber scaffold differs from Example 1 in that the amount of curcumin used is different, namely 0g, 0.03g, and 0.05g (the mass ratio of polylactic acid to curcumin is 1:0, 1:0.03, and 1:0.05, respectively).

[0037] The anisotropic drug-loaded core-shell fiber scaffolds prepared in Example 4 are designated as Product 8, Product 9, and Product 10, respectively.

[0038] Example 5 An anisotropic drug-loaded core-shell fiber scaffold differs from Example 1 in that the amount of heparin sodium used is different, namely 0g, 0.01g, and 0.05g (the mass ratio of polycaprolactone to heparin sodium is 1:0, 1:0.01, and 1:0.05, respectively).

[0039] The anisotropic drug-loaded core-shell fiber scaffolds prepared in Example 4 are referred to as Product 11, Product 12 and Product 13, respectively.

[0040] Microscopic morphology analysis was performed on products (products 1-4) prepared with different amounts of the first and second solvents to observe the morphology of the fibers in the prepared fiber scaffolds. The scanning electron microscope images are shown below. Figures 2-5 As shown, where Figure 2 For product 1 ( Figure 2 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 3 For product 2 ( Figure 3In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 4 For product 3 ( Figure 4 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 5 For product 4 ( Figure 5 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 3 , Figure 4 It can be seen that the fibers prepared under low-mass-fraction spinning solution conditions (products 2 and 3) exhibit a distinct beaded structure, indicating poor spinnability. With increasing spinning solution mass fraction, the fiber morphology gradually improves, the porous structure becomes more uniform, and the fiber diameter distribution becomes more consistent. When the spinning solution mass fraction is increased to 13 wt% (product 1, ...), the fiber morphology gradually improves, the porous structure becomes more uniform, and the fiber diameter distribution becomes more consistent. Figure 1 When the fiber is at the optimal level, the resulting fiber exhibits the most regular porous morphology and the best uniformity.

[0041] Microscopic morphology analysis was performed on products prepared with different volume percentages of anhydrous ethanol in the first and second solvents. The morphology of the fibers in the prepared fiber scaffolds was observed, and the scanning electron microscope images are shown below. Figure 2 and Figures 6-8 As shown, where Figure 2 For product 1 ( Figure 2 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 6 For product 5 ( Figure 6 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 7 For product 6 ( Figure 7 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 8 For product 7 ( Figure 8 In the diagram, 'a' and 'b' represent different magnifications of the sample, and the upper right corner of 'a' and 'b' shows the fiber diameter distribution. Figure 6 It can be seen that when the volume percentage of anhydrous ethanol is 0 (Product 5), the fiber surface still exhibits a certain porous structure, mainly due to the surface shrinkage and phase separation behavior caused by the rapid volatilization of dichloromethane. When the volume percentage of anhydrous ethanol is 0.03 (Product 1), Figure 2 The fibers exhibit optimal porous morphology, indicating that the non-solvent-induced phase separation effect is most significant under these conditions, which is beneficial for forming a uniform porous structure. However, as the volume percentage of anhydrous ethanol further increases, the overall volatility of the solvent system decreases, the phase separation kinetics slow down, resulting in a gradual decrease in both fiber porosity and pore size, and a corresponding weakening of the porous effect. When the volume percentage of anhydrous ethanol is 0.09 (product 7), Figure 8 ), the fibers tend to be non-porous.

[0042] The orientation degree of the anisotropic drug-loaded core-shell fiber scaffold was calculated using the Directionality plugin in the FIJI software. The orientation degree was defined as the percentage of the area of ​​oriented fibers (i.e., the area of ​​nanofibers arranged between 80° and 100°) in the scanning electron microscope (SEM) image relative to the total area of ​​all fibers in the SEM image. The calculated orientation degree for Product 1 was 0.519, indicating superior anisotropy.

[0043] The transmission electron microscopy (TEM) images of products 1 and 7 are shown below. Figure 9 ( Figure 9 (where a1-a4 represent different regions of the sample) and Figure 10 ( Figure 10 (A1-A4 in the diagram represent different regions of the sample). Figure 4 Both fibers exhibited a clear core-shell layered structure, and no obvious drug particles were observed. This indicates that curcumin and heparin sodium were successfully loaded onto the shell and core layers of the fibers, respectively, and that the core-shell structure achieved effective physical isolation, thereby preventing drug aggregation.

[0044] The release behavior of curcumin and heparin sodium in products 1 and 7 was determined within 0 to 14 days, and the results are as follows: Figure 11 As shown ( Figure 11 In Figure a, the release curve of curcumin is shown; in Figure b, the release curve of heparin sodium is shown. Figure 11 As can be seen, the cumulative release rate of Product 7 was 22.75% at 24 h, increasing to 63.40% at 336 h; while the release rates of Product 1 at the corresponding time points were 18.67% and 58.19%, respectively. The results indicate that the overall release rate of curcumin in Product 1 was lower than that in Product 7. This difference is closely related to the influence of fiber structure on drug distribution. During the forming process of Product 7, the polymer jet is stretched in an electric field and the solvent evaporates rapidly. Hydrophobic curcumin molecules easily migrate to the fiber surface during phase separation, forming surface enrichment, resulting in some drugs not being deeply encapsulated, thus leading to faster release. In Product 1, the non-solvent-induced phase separation process causes the drug to be encapsulated within the polymer skeleton constituting the internal pores. The drug needs to undergo a longer diffusion path to be released, therefore its initial burst release effect and overall release rate are both lower. Since curcumin has a longer duration of activity, the more sustained-release and longer-lasting release behavior provided by Product 1 helps to prolong its efficacy duration, thereby improving therapeutic effects. Figure 11As shown in b, the cumulative release rate of Product 7 was 33.46% at 24 h and 67.69% at 336 h; while the release rates of Product 1 at the corresponding time points were 39.17% and 80.11%, respectively. In contrast to the release behavior of curcumin, the release rate of heparin sodium in Product 1 was significantly higher than that in Product 7. This difference is mainly attributed to the synergistic effect of the hydrophilic properties and porous structure of heparin sodium: the well-developed internal pore structure of Product 1 greatly increases the specific surface area, promoting the penetration and contact of the release medium (water), thereby accelerating the dissolution and diffusion of hydrophilic heparin sodium. Heparin sodium has a relatively short duration of activity in vitro; the higher initial release and faster overall kinetics provided by Product 1 are beneficial for fully exerting its pharmacological activity in the short term, thereby improving therapeutic efficacy.

[0045] To investigate the effect of drug-loaded scaffolds on cell proliferation, human adipose-derived mesenchymal stem cells were seeded onto products with different drug loading amounts. Cell viability was assessed using the CCK-8 assay on days 7 and 14. The results are as follows: Figure 12 As shown ( Figure 12 In Figure a, cell viability on days 7 and 14 is shown for products with different curcumin dosages; in Figure b, cell viability on days 7 and 14 is shown for products with different heparin sodium dosages. The results showed that product 1, loaded with 1% curcumin (polylactic acid to curcumin mass ratio 1:0.01) and 3% heparin sodium (polycaprolactone to heparin sodium mass ratio 1:0.03), exhibited significantly better cell viability (absorbance) than other products, indicating that this drug combination most effectively promotes cell proliferation. This result demonstrates that the core-shell structure and porous morphology synergistically optimize the sustained-release behavior of the drug, thereby significantly improving the biocompatibility of the scaffold and providing a more favorable microenvironment for cell growth and proliferation.

[0046] Based on the above experimental verification, the porous core-shell fiber scaffold prepared in this application achieves independent regulation of the partitioned loading and release behavior of curcumin and heparin sodium through a unique core-shell structure design and component matching. The principle of partitioned loading is as follows: the two drugs are dissolved separately in different polymer spinning systems, and coaxial electrospinning is used to form core-shell fibers with a stable interface, thereby achieving physical isolation of the drugs. This structure effectively avoids compatibility problems that may be caused by the differences in physicochemical properties between curcumin and heparin sodium, ensuring that the two do not interfere with each other during storage and release.

[0047] In terms of controlled drug release, the core-shell structure and the porous surface morphology work synergistically to construct two distinct drug release kinetics. For curcumin in the first shell, the porous structure embeds it deeper within the polylactic acid backbone, extending the diffusion path and achieving long-lasting sustained release. For heparin sodium in the core, the porous structure significantly increases the fiber surface area, promoting rapid penetration of the release medium and drug diffusion, resulting in a significantly improved release rate. Under specific conditions, the release kinetics of the two types of drugs can be independently and precisely controlled by spinning process parameters such as drug loading and the core-shell propulsion speed ratio, meeting the drug therapy needs in complex treatment scenarios.

[0048] In the description of this specification, the terms "first" and "second" are used for descriptive purposes only and should not be construed as indicating or implying relative importance or implicitly specifying the number of indicated technical features. Thus, a feature defined as "first" or "second" may explicitly or implicitly include at least one of that feature.

[0049] The above description is only a preferred embodiment of the present invention. It should be noted that those skilled in the art can make several improvements and additions without departing from the method of the present invention, and these improvements and additions should also be considered within the scope of protection of the present invention.

Claims

1. An anisotropic drug-loaded core-shell fiber scaffold, characterized in that, It consists of a core, a first shell, and a second shell, in sequence. The core layer includes a porous core layer matrix and a hydrophilic drug loaded on the porous core layer matrix, wherein the porous core layer matrix is ​​formed of polycaprolactone; The first shell layer includes a porous shell matrix and a hydrophobic drug loaded on the porous shell matrix, wherein the porous shell matrix is ​​formed of polylactic acid; The second shell layer is obtained by cross-linking and curing modified gelatin with genipin, wherein the modified gelatin is obtained by methacrylating gelatin.

2. The anisotropic drug-loaded core-shell fiber scaffold according to claim 1, characterized in that, The method for preparing the modified gelatin includes the following steps: The gelatin was dissolved in a carbonate buffer solution, the pH was adjusted to 8-10, and methacrylic anhydride was added under the condition of heating to 40℃-60℃. After stirring and reacting, the pH was adjusted to 6-8, dialyzed, and freeze-dried to obtain the modified gelatin. The mass-to-volume ratio of the gelatin to the methacrylic anhydride is 10 g:(0.5-2) mL.

3. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to any one of claims 1-2, characterized in that, The modified gelatin solution is prepared from the modified gelatin, and the genipin solution is prepared from the genipin; the preparation method of the anisotropic drug-loaded core-shell fiber scaffold includes the following steps: The polylactic acid and the hydrophobic drug are added to a first solvent to prepare spinning solution A; The polycaprolactone and the hydrophilic drug are added to a second solvent to prepare spinning solution B; The spinning solution A and the spinning solution B are respectively loaded into a syringe and coaxial electrospinning is performed to obtain an electrospun material. First, the electrospun material is placed in a Tween 80 solution for hydrophilization treatment, then placed in a modified gelatin solution for adsorption of the modified gelatin, and finally placed in a genipin solution for cross-linking and curing to obtain the anisotropic drug-loaded core-shell fiber scaffold. The first solvent comprises dichloromethane, hexafluoroisopropanol and anhydrous ethanol in a volume ratio of 0.85:0.15:(0-0.06), and the second solvent comprises dichloromethane, hexafluoroisopropanol and anhydrous ethanol in a volume ratio of 0.85:0.15:(0-0.06).

4. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The spinning solution A has a mass fraction of 11wt%-15wt%, and the spinning solution B has a mass fraction of 11wt%-15wt%.

5. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The propulsion speed of spinning solution A is 0.5 mL / h-2 mL / h, and the propulsion speed of spinning solution B is 0.1 mL / h-0.5 mL / h.

6. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The mass ratio of polylactic acid to the hydrophobic drug is 1:(0.01-0.05), and the mass ratio of polycaprolactone to the hydrophilic drug is 1:(0.01-0.05).

7. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The Tween 80 solution has a mass fraction of 0.05wt%-0.2wt%, and the hydrophilization treatment takes 30min-60min.

8. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The modified gelatin solution has a mass fraction of 1wt%-5wt%, and the adsorption time is 30min-120min.

9. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The mass fraction of the genipin solution is 0.05wt%-0.2wt%, and the cross-linking curing time is 2h.

10. The method for preparing the anisotropic drug-loaded core-shell fiber scaffold according to claim 3, characterized in that, The coaxial electrospinning time is 180 min-360 min, and the coaxial electrospinning collection speed is 1000 rpm-2000 rpm.

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