Programmable drug delivery system responding to proteome signal and preparation method thereof

Abstract

The invention relates to the technical field of biological materials, and provides a programmable drug delivery system responding to proteome signals and a preparation method thereof. The carrier comprises a degradable polymer skeleton and a responsive functional layer on the surface of the degradable polymer skeleton, the functional layer contains a matrix capable of being degraded by matrix metalloproteinase, a first therapeutic agent and a second therapeutic agent, and the two therapeutic agents are spatially separated through core-shell structure design. When implanted into a wound surface, the matrix is subjected to enzymolysis, the first therapeutic agent is triggered to be rapidly released preferentially to deal with early risks, and then the second therapeutic agent is continuously released to promote later repair, so that sequential and programmed intelligent administration is realized. The preparation method is particularly suitable for preparing medical instruments such as intelligent sutures capable of promoting healing of complicated wounds such as diabetic foot ulcers and chronic infected wounds, and the problems that a traditional medicine carrying instrument is disordered in release and disjointed with the healing stage are solved.

Description

Technical Field

[0001] This invention relates to the field of biomaterials technology, specifically to a programmable drug delivery system that responds to proteome signals and its preparation method. Background Technology

[0002] Modern surgery is constantly evolving towards minimally invasive and precision medicine, and the demand for sutures has far exceeded the scope of traditional mechanical fixation. However, the current clinical situation has significant shortcomings: the widely used absorbable sutures have limited functions, cannot provide active treatment during the healing process, and their degradation process may produce acidic byproducts, increasing the risk of local inflammation. Although drug-loaded sutures have been developed, their drug release mechanism still relies on passive diffusion or material hydrolysis, and cannot sense the biological signals of dynamic changes in the wound microenvironment. This leads to a serious disconnect between drug release and the actual needs of the healing stage, often resulting in early burst drug release followed by insufficient follow-up, making it difficult to meet the clinical needs of complex wound time-sequential and individualized management.

[0003] The key to overcoming existing technological bottlenecks lies in achieving "intelligent" sutures, that is, sutures capable of actively sensing and responding to specific physiological signals of the wound. Modern molecular biology has revealed that specific proteomic signals in the wound microenvironment (such as overexpressed matrix metalloproteinases, MMPs) are key biomarkers characterizing the healing state. Therefore, developing a suture capable of directly "interpreting" these proteomic signals and executing programmable drug release logic on demand is an inevitable trend from static suturing to dynamic precision medical management, and also a fundamental way to overcome the limitations of existing functional sutures in terms of release accuracy, functional integration, and material biocompatibility. Therefore, there is an urgent need to develop an intelligent drug delivery system capable of actively sensing wound protease signals and achieving the sequential and programmed release of multiple therapeutic agents through specific spatial structural design to meet the dynamic needs of complex wound healing. Summary of the Invention

[0004] In view of the problems of passive and disordered drug release in existing drug-loaded medical devices (such as sutures) and mismatch with the dynamic needs of complex wound healing processes, this invention aims to provide an intelligent drug delivery system that can actively sense key biological signals (proteomic signals) in the wound microenvironment and thereby realize the sequential and programmed release of various therapeutic agents.

[0005] To address the aforementioned problems, the present invention provides a programmable drug delivery system responsive to proteomic signals, comprising: a degradable polymer backbone, and a responsive functional layer constructed on its surface; the responsive functional layer includes a matrix that can be degraded by matrix metalloproteinases, a first therapeutic agent, and a second therapeutic agent, wherein the first therapeutic agent and the second therapeutic agent are spatially separated in the functional layer through a core-shell structure to achieve the sequential release of the two therapeutic agents triggered by the matrix metalloproteinases.

[0006] The system comprises a degradable polymer backbone and a responsive functional layer encapsulated on its surface; the responsive functional layer comprises a matrix that can be degraded by matrix metalloproteinases in the wound microenvironment, a first therapeutic agent, and a second therapeutic agent; the functional layer includes a core-shell structure, wherein the matrix forms a shell, the second therapeutic agent is encapsulated within the core, and the first therapeutic agent is loaded in the shell matrix; when the matrix metalloproteinases degrade the matrix, the first therapeutic agent is released preferentially over the second therapeutic agent.

[0007] Preferably, the biodegradable polymer backbone is at least one of polylactic acid-glycolic acid copolymer (PLGA), polycaprolactone (PCL), polylactic acid (PLA), or copolymers or mixtures thereof; the polymer backbone is fiber or yarn.

[0008] Preferably, the matrix metalloproteinase is a gelatinase overexpressed in inflammatory or abnormally healing wounds. More preferably, the gelatinase is at least one of gelatinase A (MMP-2) and gelatinase B (MMP-9). The matrix comprises a biomaterial that can be specifically degraded by the aforementioned MMPs, such as gelatin, collagen, gelatin-polysaccharide complexes, or derivatives thereof.

[0009] Preferably, the first therapeutic agent is a drug used to control early symptoms, such as antibiotics (levofloxacin, moxifloxacin, etc.) or nonsteroidal anti-inflammatory drugs (such as ketoprofen). The second therapeutic agent is a drug used to promote medium- to long-term healing, such as growth factors (vascular endothelial growth factor VEGF, basic fibroblast growth factor bFGF, platelet-derived growth factor PDGF, etc.) or cytokines (such as interleukin-10).

[0010] Preferably, the responsive functional layer is constructed on the surface of the polymer skeleton by coaxial electrospinning, coaxial microfluidics, or layer-by-layer self-assembly technology, and preferably forms a core-shell structured nanofiber coating or microfiber coating.

[0011] Secondly, the present invention provides a specific product comprising the above-described system—a programmable drug delivery suture responsive to proteomic signals. The polymer backbone of the suture is a surgical absorbable suture (such as PLGA absorbable yarn), and its surface is constructed with the aforementioned responsive functional layer.

[0012] Furthermore, using PLGA yarn as a skeleton, a gelatin enzyme-responsive functional layer was constructed by coaxial electrospinning. This functional layer contains gelatin (shell matrix), levofloxacin (loaded in gelatin), and vascular endothelial growth factor (VEGF) (located in the core layer).

[0013] Thirdly, the present invention provides a method for preparing the above-mentioned system, comprising the following steps:

[0014] S1. Prepare a shell solution (solution A) and a core solution (solution B); solution A contains an enzymatically hydrolyzable matrix and a first therapeutic agent; solution B contains a second therapeutic agent.

[0015] S2. Using a polymer skeleton as a receiving device (such as wound on a receiving roller) or as the core material for coaxial spinning, coaxial electrospinning technology is used to simultaneously spin out liquid A and liquid B, and a responsive functional layer with a core-shell structure is deposited on the surface of the polymer skeleton.

[0016] S3. Perform chemical or physical cross-linking treatment on the deposited functional layer to improve its stability and prevent non-specific disintegration during storage or early use.

[0017] Preferably, the concentration of the matrix in the shell solution is 5-15 wt%, the concentration of the first therapeutic agent is 1-10 mg / mL, and the concentration of the second therapeutic agent in the core solution is 0.05-0.5 mg / mL.

[0018] Preferably, the parameters for the coaxial electrospinning are: voltage 15 kV, shell flow rate 0.5-1.5 mL / h, core flow rate 0.2-0.8 mL / h, spinning distance 15 cm, ambient temperature 25 ℃, and relative humidity 40%.

[0019] Preferably, the crosslinking treatment uses a biocompatible crosslinking agent, such as genipin (concentration 0.1-1 wt%), and the crosslinking time is 6-12 hours.

[0020] Fourthly, the present invention provides the use of the above-described system or suture in the preparation of medical devices for promoting wound healing. The wounds are characterized by the overexpression of matrix metalloproteinases (such as MMPs) and are often accompanied by infection, chronic inflammation, or healing impairment, such as diabetic foot ulcers, pressure sores, venous ulcers, gastrointestinal surgical anastomoses, deep burn wounds, etc.

[0021] The advantages of this invention are:

[0022] 1. This invention provides a release mechanism using matrix metalloproteinases in the wound microenvironment as biological signal triggers, constructing an intelligent closed loop of "signal sensing-system response-programmed drug release". Through the core core-shell structure design, programmable drug delivery is precisely achieved by "rapidly releasing antibacterial drugs first, followed by sustained release of growth factors", which helps to solve the problems of passive, disordered drug release and serious disconnection from the healing stage in traditional drug-loaded sutures.

[0023] 2. This invention ingeniously constructs a composite material system of "PLGA skeleton-gelatin functional layer", which is not only safe and biodegradable in all components and does not require secondary suture removal, but also significantly reduces the risk of inflammation caused by the material itself while providing sufficient mechanical support, thus achieving a unity of device function and tissue friendliness.

[0024] 3. This invention upgrades sutures from a single mechanical fixation tool to a comprehensive management platform that can dynamically sense wound status and actively intervene in treatment. Through the sequential synergy of anti-infection and repair-promoting functions, it significantly reduces the risk of postoperative complications. Attached Figure Description

[0025] Figure 1 The images shown are (a) scanning electron microscope (SEM) and (b) transmission electron microscope (TEM) images of the programmable drug delivery suture prepared in Example 1 of this invention, showing the surface nanofibers and core-shell structure.

[0026] Figure 2 The graph shows a comparison of the mechanical properties of the samples prepared in Example 1 and Comparative Examples 1 and 3.

[0027] Figure 3 The in vitro degradation curves of the samples prepared in Example 1 and Comparative Examples 1-3 in (a) PBS solution and (b) PBS solution containing gelatinase are shown.

[0028] Figure 4 The cumulative release curves of (a) levofloxacin and (b) VEGF in gelatinase-containing media prepared in Examples 1 and Comparative Examples 1-3 demonstrate the time-sequential release characteristics of the present invention.

[0029] Figure 5 The following is an evaluation diagram of the in vivo healing effect of the suture prepared in Example 1 in a rat model of infectious skin defects: Figure 5 (ah) represents the infiltration of inflammatory cells, growth of granulation tissue, and angiogenesis in rat skin.

[0030] Figure 6 In vivo biocompatibility evaluation of the sutures prepared in Example 1 (H&E staining of major organs). Detailed Implementation

[0031] The technical solutions described in this invention will now be clearly and completely described with reference to the accompanying drawings in the embodiments of this invention.

[0032] It is understood that the core of this invention lies in a universal platform for time-sequential release through an enzymatically hydrolyzable matrix and a core-shell structure. The first therapeutic agent is not limited to levofloxacin; any drug suitable for controlling early wound infection, inflammation, or oxidative stress, such as other quinolone antibiotics (e.g., moxifloxacin), aminoglycoside antibiotics (e.g., gentamicin), and nonsteroidal anti-inflammatory drugs (e.g., ketoprofen), can be loaded into the shell matrix in a similar manner for release in the first stage.

[0033] The second therapeutic agent is not limited to VEGF. Any bioactive factor designed to promote tissue regeneration and repair, such as other growth factors (e.g., bFGF, PDGF) or cytokines with anti-inflammatory or immunomodulatory effects (e.g., interleukin-10), can be encapsulated in the nuclear layer. As long as its activity is maintained in the selected solvent and preparation process, it can be released in the second stage after matrix degradation.

[0034] The matrix is ​​not limited to gelatin. Any biocompatible material that can be specifically degraded by matrix metalloproteinases (preferably at least one of gelatinases MMP-2 and MMP-9) overexpressed in the target wound, such as collagen from different sources, modified gelatin derivatives or their complexes with polysaccharides, can be used as a responsive matrix.

[0035] Those skilled in the art can select matching matrix materials based on the protease expression profile of the target wound; select appropriate combinations of first and second therapeutic agents according to the needs of the pathological stage of the wound; and refer to the preparation method of Example 1 to achieve the time-release effect of the present invention without creative labor.

[0036] The biodegradable polymer framework of the present invention is not limited to the fiber or yarn form shown in the embodiments, but can also be other forms suitable as drug delivery systems such as films, mesh scaffolds, and microparticles. The responsive functional layer can be coated or constructed on the surface of the framework by techniques including but not limited to coaxial electrospinning, layer-by-layer self-assembly, and microfluidic coating to form structures such as nano / micro fiber coatings.

[0037] Obviously, the embodiments described in this specification are only some feasible technical solutions of the present invention. Other implementation methods obtained by those skilled in the art based on the embodiments of the present invention without any creative effort should be considered to fall within the scope of protection of the present invention.

[0038] Example 1: Preparation of a programmable drug delivery suture responsive to gelatinase

[0039] This embodiment prepares a specific suture product.

[0040] 1. Solution preparation:

[0041] (1) Shell solution A: Weigh gelatin and dissolve it in an acetic acid-water mixed solvent (volume ratio 7:3) to prepare a 10 wt% solution. After stirring to dissolve, add levofloxacin powder to make a final concentration of 5 mg / mL, and continue stirring until completely dissolved / dispersed.

[0042] (2) Core layer solution B: Dissolve VEGF in phosphate buffer (PBS) at pH 7.4 to prepare a solution with a concentration of 0.1 mg / mL.

[0043] 2. Coaxial electrospinning

[0044] Commercially available PLGA absorbable suture with a diameter of 0.2 mm was selected as the receiving skeleton and tightly wound onto a grounded receiving roller. Using a coaxial spinning needle, shell solution A and core solution B were injected into the syringes on the outer (shell) and inner (core) axes, respectively. Specific spinning parameters were set as follows: voltage 15 kV, receiving roller grounded, shell flow rate 1.0 mL / h, core flow rate 0.5 mL / h, distance from needle tip to receiving skeleton (spinning distance) 15 cm, ambient temperature 25℃, and relative humidity 40%. The high-voltage power supply and injection pump were turned on, and spinning was continued on the PLGA suture surface for 2 hours to form a uniform nanofiber functional coating.

[0045] 3. Crosslinking treatment

[0046] After spinning, the sutures loaded with the functional coating were immersed in a 0.5 wt% genipin aqueous solution and crosslinked at 4°C for 6 hours. After crosslinking, the sutures were thoroughly washed with deionized water to remove unreacted genipin, and then freeze-dried to obtain the final product.

[0047] Comparative Example 1: Preparation of Physically Mixed Drug-Loaded Sutures

[0048] This comparative example simulates a traditional mixed drug delivery method. The difference from Example 1 is the use of uniaxial electrospinning. The spinning solution was prepared by physically mixing levofloxacin, VEGF, and PLGA in a mixed solvent of dichloromethane (DCM) and N,N-dimethylformamide (DMF). Using PLGA sutures as the receiving substrate, a mixed drug-loaded fiber layer was constructed using this mixed spinning solution via uniaxial electrospinning (voltage 15 kV, flow rate 1.0 mL / h, distance 15 cm). After spinning for the same time, the fiber was directly dried without genipin crosslinking. In this sample, the two drugs were randomly distributed within the PLGA fibers. The concentration of levofloxacin was 5 mg / mL, the concentration of VEGF was 0.1 mg / mL, the amount of PLGA used was 10 wt% of the spinning solution, and the volume ratio of dichloromethane (DCM) to N,N-dimethylformamide (DMF) in the mixed solvent was 7:3.

[0049] Comparative Example 2: Preparation of Enzyme-Independent Double-Layer Drug-Loaded Sutures

[0050] This comparative example simulates a time-sequential release attempt without specific responsiveness. The difference from Example 1 is that the gelatin in shell solution A is replaced with a non-gelatinase-responsive polymer, such as PLGA (10 wt%, dissolved in DCM:DMF=7:3), still containing levofloxacin (5 mg / mL); core solution B remains unchanged (VEGF, 0.1 mg / mL). The exact same coaxial electrospinning and crosslinking parameters as in Example 1 are used. This sample has a core-shell structure, but its degradation depends on the hydrolysis of PLGA, not an enzymatic response, and it is not crosslinked.

[0051] Comparative Example 3: Preparation of a hybrid drug-loaded suture material with enzyme responsiveness but no core-shell structure

[0052] This comparative example simulates a case where only responsive materials are used without a core structural design. The difference from Example 1 is the use of uniaxial electrospinning. The spinning solution was prepared by physically mixing levofloxacin, VEGF, and PLGA in a mixed solvent of dichloromethane (DCM) and N,N-dimethylformamide (DMF). Using PLGA sutures as the receiving substrate, a mixed drug-loaded fiber layer was constructed using this mixed spinning solution via uniaxial electrospinning (voltage 15 kV, flow rate 1.0 mL / h, distance 15 cm). After the same spinning time, the sample was crosslinked with 0.5 wt% genipin for 6 hours. Although gelatin was used in this sample, both drugs were mixed within the gelatin matrix without spatial separation. The concentration of levofloxacin is 5 mg / mL, the concentration of VEGF is 0.1 mg / mL, the amount of PLGA used is 10 wt% of the spinning solution, and the volume ratio of dichloromethane (DCM) to N,N-dimethylformamide (DMF) in the mixed solvent is 7:3.

[0053] Example 2: Preparation of a programmable drug delivery suture responsive to gelatinase

[0054] The difference compared to Example 1 is as follows:

[0055] 1. Solution preparation: Replace gelatin with gelatin-polysaccharide complex to make the collagen concentration 5 wt%, replace levofloxacin with ketoprofen to make the final concentration of ketoprofen 1 mg / mL, and replace VEGF with interleukin-10 to make the final concentration of interleukin-10 0.05 mg / mL.

[0056] 2. Coaxial electrospinning: Replace PLGA with PLA. The coaxial electrospinning parameters are: voltage 15 kV, shell flow rate 0.5 mL / h, core flow rate 0.2 mL / h, spinning distance 15 cm, ambient temperature 25 ℃, and relative humidity 40%.

[0057] 3. Crosslinking treatment: Immerse in a 0.1 wt% genipin solution for 12 h for crosslinking.

[0058] Example 3: A programmable drug delivery patch responsive to MMP

[0059] This embodiment illustrates that the drug delivery system of the present invention is not limited to the suture morphology.

[0060] 1. Solution preparation: Compared with Example 1, the difference is that gelatin is replaced with collagen to make the collagen concentration 15wt%, levofloxacin is replaced with gentamicin to make the final concentration of gentamicin 10 mg / mL, and VEGF is replaced with PDGF to make the final concentration of PDGF 0.5 mg / mL.

[0061] 2. Preparation: A flat aluminum foil was used as the receiving plate. The coaxial electrospinning parameters were: voltage 15 kV, shell flow rate 1.5 mL / h, core flow rate 0.8 mL / h, spinning distance 15 cm, ambient temperature 25 ℃, and relative humidity 40%. Nonwoven film was formed by spinning directly on the receiving plate.

[0062] 3. Post-processing: The spun film is peeled off from the aluminum foil, immersed in genipin solution (concentration of 1 wt%) for cross-linking for 12 hours, washed and dried to obtain a biodegradable patch with time-release drug function.

[0063] The resulting patch, whose functional layer consists of nanofibers with the same chemical composition, microstructure, and cross-linking state as the suture surface coating described in Example 1, is therefore expected to exhibit similar enzyme-responsive degradation behavior and a sequential drug release characteristic of "first therapeutic agent, then second therapeutic agent" when applied to wounds containing matrix metalloproteinases (such as MMP-2 / 9). This embodiment demonstrates the feasibility of extending the core intelligent delivery system of this invention from a fiber morphology (suture) to a two-dimensional film morphology (patch / dressing).

[0064] Experimental Example 1: Physical characterization of the programmable drug delivery suture prepared in Example 1

[0065] The microstructure of the sutures prepared in Example 1 was characterized using scanning electron microscopy (SEM) and transmission scanning electron microscopy (TEM). The results are as follows: Figure 1 As shown in (ab), there are a large number of nanofibers of about 500 nm on the surface of the suture, and the nanofibers can be seen to have a distinct core-shell structure.

[0066] Experimental Example 2: Mechanical property characterization of the programmable drug delivery suture prepared in Example 1

[0067] The mechanical properties of the sutures prepared in Example 1 were evaluated using a universal testing machine. Suture samples from Example 1, Comparative Example 1, and Comparative Example 3 (n=5) were cut into 10 cm lengths and tested at a tensile rate of 10 mm / min, with the breaking strength recorded. The results are as follows: Figure 2 As shown, the breaking strength of Example 1 and Comparative Examples 1 and 3 are all higher than 8 N. This indicates that whether PLGA or gelatin is used as the functional layer, it can provide sufficient mechanical support for the suture without sacrificing its core mechanical properties due to drug loading or structural design, thus meeting the basic requirements of surgical suturing. In summary, the suture prepared in Example 1 has good mechanical properties.

[0068] Experimental Example 3: Characterization of the degradation behavior of the programmable drug delivery suture prepared in Example 1

[0069] To evaluate the degradation characteristics of the sutures prepared in Example 1 under different environments, suture samples (n=3) from Example 1, Comparative Example 1, Comparative Example 2, and Comparative Example 3 were immersed in PBS buffer at pH 7.4 and PBS solution containing 300 μg / mL gelatinase, respectively, and placed in a shaker at 37 ℃. Samples were taken periodically, dried, and weighed. The weight loss rate was calculated for the PBS solution containing gelatinase and the PBS solution. The results are shown below. Figure 3As shown in (a) and (b), the experimental groups degraded slowly in PBS; however, in a gelatinase environment, the gelatin functional layers of Example 1 and Comparative Example 3 were rapidly enzymatically hydrolyzed within 6 h, resulting in a rapid decline in quality. In contrast, the PLGA fiber layers of Comparative Example 1 and Comparative Example 2 exhibited slow hydrolysis characteristics. This indicates that the sutures prepared in Example 1 are gelatinase responsive and can respond to proteomic signals.

[0070] Experimental Example 4: Characterization of drug release behavior of the programmable drug delivery suture prepared in Example 1

[0071] To evaluate the intelligent response and time-sequential release capability of the sutures prepared in Example 1, suture samples (n=3) from Example 1, Comparative Example 1, Comparative Example 2, and Comparative Example 3 were placed in PBS release medium containing gelatinase (300 μg / mL) and shaken at 37 °C. 3 mL of the test solution was collected at time points 1, 6, 12, 24, and 48 h, and an equal volume of fresh medium was added simultaneously. The concentration of levofloxacin was detected by HPLC, and the concentration of VEGF was detected by ELISA; the cumulative release rate was calculated accordingly.

[0072] The cumulative release rates of levofloxacin and VEGF were as follows: Figure 4 As shown in (a) and (b), in Example 1, under gelatinase triggering, levofloxacin achieved a release rate of 82% within 24 hours, followed by sustained release of VEGF. In contrast, Comparative Examples 1 and 3 exhibited both rapid burst releases and disordered synchronous releases of the two drugs, making it impossible to distinguish the stages. In Comparative Example 2, due to the blocking effect of PLGA, the release of both drugs was extremely slow, with only 33% of levofloxacin and 27% of VEGF released within 24 hours. In summary, the suture prepared in Example 1 possesses programmed drug release capability.

[0073] Experimental Example 5: Characterization of the in vivo anti-inflammatory and healing effects of the programmable drug delivery suture prepared in Example 1.

[0074] To evaluate the in vivo anti-inflammatory and healing effects of the sutures prepared in Example 1, a rat model of infectious skin defects was established. Rats were randomly divided into the Example 1 group and a blank control group, and the corresponding sutures were implanted. Postoperatively, wound infection and healing rates were observed and recorded regularly. Histological analysis was performed on days 7 and 14 to assess inflammatory cell infiltration, granulation tissue growth, and angiogenesis.

[0075] The results are as follows Figure 5 As shown, Example 1 group exhibited the strongest infection control capability in the early stages. Figure 5 ef), and exhibits the richest angiogenesis and fastest wound closure in the later stages (ef), Figure 5 In contrast, the control group showed insufficient anti-infection and repair-promoting effects (gh). Figure 5 (ad). In summary, the sutures prepared in Example 1 have good anti-infection and healing capabilities.

[0076] Experimental Example 6: In vivo biosafety characterization of the programmable drug delivery suture prepared in Example 1

[0077] To assess the biocompatibility of the sutures prepared in Example 1, we euthanized mice on the last day of Experiment 5 and harvested vital organs (spleen, liver, heart, kidney, lung, and brain). After H&E staining, the vital organs were fixed, embedded, and subjected to histological examination. In any treatment group, such as... Figure 6 As shown, no necrosis or significant tissue damage (including apoptosis, nuclear cleavage, pyknosis, or inflammatory cell infiltration) was detected in the Example 1 group compared to the control group. This indicates that the dosage used was safe and had no obvious toxicity.

[0078] 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 principle of the present invention, and these improvements and additions should also be considered within the scope of protection of the present invention.

Claims

1. A programmable drug delivery system responsive to proteomic signals, characterized in that, The system comprises a degradable polymer skeleton and a responsive functional layer coated on the surface of the skeleton; the responsive functional layer comprises a matrix that can be degraded by matrix metalloproteinase in the microenvironment of a wound surface, a first therapeutic agent and a second therapeutic agent; the functional layer comprises a core-shell structure, wherein the matrix forms a shell layer, the second therapeutic agent is coated in a core layer, and the first therapeutic agent is loaded in the shell layer matrix; when the matrix is degraded by the matrix metalloproteinase, the first therapeutic agent is triggered to be released preferentially to the second therapeutic agent. The system comprises a degradable polymer skeleton and a responsive functional layer coated on the surface of the skeleton; the responsive functional layer comprises a matrix that can be degraded by matrix metalloproteinase in the microenvironment of a wound surface, a first therapeutic agent and a second therapeutic agent; the functional layer comprises a core-shell structure, wherein the matrix forms a shell layer, the second therapeutic agent is coated in a core layer, and the first therapeutic agent is loaded in the shell layer matrix; when the matrix is degraded by the matrix metalloproteinase, the first therapeutic agent is triggered to be released preferentially to the second therapeutic agent.

2. The system of claim 1, wherein, The degradable polymer skeleton is at least one of polylactic acid-glycolic acid copolymer, polycaprolactone, polylactic acid or a copolymer thereof.

3. The system of claim 1, wherein, The matrix metalloproteinase is any one of MMP-2 and MMP-9; the matrix comprises any one of gelatin, collagen, gelatin-polysaccharide complex or a derivative thereof.

4. The system of claim 1, wherein, The first therapeutic agent is an antibiotic or a non-steroidal anti-inflammatory drug; the second therapeutic agent is a growth factor or a cytokine.

5. The system of claim 1, wherein, The antibiotic is any one of levofloxacin, moxifloxacin or gentamicin; the non-steroidal anti-inflammatory drug is ketoprofen; the growth factor is any one of vascular endothelial growth factor, basic fibroblast growth factor or platelet-derived growth factor; the cytokine is interleukin 10.

6. The system of claim 5, wherein, The polymer skeleton is an absorbable suture, and the absorbable suture is coated with the responsive functional layer according to any one of claims 1-6.

7. A programmable drug delivery suture in response to a proteomic signal according to any one of claims 1-6, characterized in that, The absorbable suture is a PLGA absorbable yarn; and the responsive functional layer is a gelatinase-responsive functional layer comprising gelatin, levofloxacin and vascular endothelial growth factor.

8. The suture of claim 7, wherein, The method comprises the following steps:

9. A method of preparing a system according to any one of claims 1-6, characterized in that, S1. Preparing a shell layer solution comprising the matrix and the first therapeutic agent, and a core layer solution comprising the second therapeutic agent; S2. Using the polymer skeleton as a receiving substrate or core material, using the shell layer solution of S1 as a shell layer spinning solution and the core layer solution of S1 as a core layer spinning solution, and using a coaxial electrospinning technology to deposit the responsive functional layer with a core-shell structure on the surface of the polymer skeleton; The parameters of the coaxial electrospinning are as follows: voltage 15 kV, shell layer flow rate 0.5-1.5 mL / h, core layer flow rate 0.2-0.8 mL / h, spinning distance 15 cm, ambient temperature 25 ℃ and relative humidity 40%; S3. Crosslinking the deposited functional layer to stabilize its structure; The crosslinking treatment uses a crosslinking agent, genipin, and the concentration of the genipin is 0.1-1 wt%, and the crosslinking treatment time is 6-12 h. ​ 10. The preparation method according to claim 9, characterized in that, The concentration of the matrix in the shell solution is 5-15 wt%, and the concentration of the first therapeutic agent is 1-10 mg / mL; the concentration of the second therapeutic agent in the core solution is 0.05-0.5 mg / mL.

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