Coating for Medical Devices

Abstract

The present invention discloses a coating for medical devices. This coating includes a drug-carrying layer and a drug-controlled release layer coated outside the drug-carrying layer. The molecular weight retention rate m of the polymer in the drug-controlled release layer, the weight-average molecular weight M W制御放出 , and the thickness X of the polymer 制御放出 satisfy the following relational expressions: JPEG2025524216000005.jpg23145Here, A1, A2, k, and b are all fitting constants. The design of this coating can not only improve the drug release performance but also effectively control the corrosion of the substrate. Therefore, the substrate of the degradable medical device will not decompose within 1 to 6 months or will decompose slowly, and will decompose rapidly after 6 months.

Description

Technical Field

[0001] The present invention belongs to the technical field of medical devices, and more specifically, relates to coatings for medical devices, and even more specifically, coatings for degradable medical devices.

Background Art

[0002] Medical devices, especially implantable medical devices and interventional medical devices, not only have basic treatment functions, but also need to carry drugs on their surfaces to prevent rejection reactions, proliferation, thrombosis, inflammation, and other reactions at the implantation sites of the devices. Therefore, currently common medical devices usually add some kind of coating to their surfaces to carry drugs or control drug release. After being implanted into blood vessels, vascular stents, for example, not only provide sufficient support for a certain period to expand blood vessels and maintain the open state of blood vessels, but also need to prevent blood vessel proliferation and problems such as thrombosis and subsequent restenosis at the implantation site. Therefore, currently commercially available vascular stents usually add one or more coatings on their surfaces to carry drugs or control drug release in order to prevent the basic treatment effect from being adversely affected by abnormal reactions after the stent is implanted into the human body.

[0003] The coating on the surface of the device has very strict requirements regarding drug loading and drug release. It is necessary to control both the drug loading amount and the drug release rate. If the drug loading amount is too high, excessive drugs may be absorbed by the human body, causing serious toxicity and side effects, increasing the risk of drug toxicity in the body. If the drug loading amount is too low, a good preventive and therapeutic effect cannot be achieved. When the total amount of the drug in the medical device is constant, if the drug release rate is too slow, the drug will be insufficient throughout the drug release stage, and the preventive and therapeutic effect of the drug will be limited. If the drug release rate is too fast, the local concentration of the drug will be too high, resulting in toxicity and side effects. Moreover, in the initial stage, the drug release is too fast, and the drug cannot be released effectively and continuously. In the later stage, the amount of the drug released will be insufficient. In any of these cases, growth occurs at the implantation site, ultimately causing vascular restenosis. Therefore, it has been found that the drug loading amount and the controlled release of drugs on the surface of medical devices are important issues in the design of drug coatings. In the case of degradable medical devices, the coating not only needs to carry drugs and control drug release but also needs to play a role in controlling the corrosion of the device substrate. Therefore, the design of the coating for degradable devices is a major issue and challenge.

[0004] At present, many research results on drug coatings have been published. However, most of the existing coatings for medical devices only consider drug release and do not take into account the impact of the coating on the corrosion of the device substrate. Therefore, the currently disclosed drug coating designs are mostly not suitable for use with degradable medical devices. On the one hand, among the coatings currently in common use, there are some that are not suitable in terms of drug release and have some defects. On the other hand, when considering the corrosion requirements of degradable medical devices in the coating design, the types and molecular weights of the polymers in the coating change greatly. For controlled drug release, different types and molecular weights of polymers will result in significant differences in controlled drug release, and the overall coating design will also change greatly. For example, Patent CN107496996B discloses a drug coating for a vascular stent. This coating includes three layers, each layer contains a drug and a drug carrier. The outer layer coating serves to control the release of the inner layer coating, thereby reducing the burst release of the drug in the inner layer and reducing the possibility of acute toxicity reactions. However, two types of drugs are carried on the outer layer of this drug coating, and the PLGA carrying the drug has a relatively small molecular weight, only about 2000. Also, the mass fraction ratio of rapamycin, curcumin, and the PLGA carrier carried on this outer layer coating is 5:2:13. The molecular weight of the polymer in this outer layer coating is small, the proportion of the polymer is low, and the proportion of the drug is high. The overall coating design does not have a good sustained release effect on the drugs in the outer layer, and there is also a phenomenon of burst release for the large amount of rapamycin carried on the outer layer. A large amount of rapamycin is carried on both the middle layer and the outer layer, and the drug loading amount far exceeds the required amount, which may cause additional toxicity and side effects to the human body. Furthermore, since the substrate of this vascular stent is stainless steel, the drug coating design in this patent application does not need to consider the degradation of the substrate. Therefore, the polymer used is a low molecular weight polymer and is not suitable for medical devices with degradable substrates.

Summary of the Invention

Problems to be Solved by the Invention

[0005] Based on this, the present application provides a coating for medical devices, particularly for degradable medical devices, which can ensure that the drug loading amount is within an appropriate effective range through a unique coating design, exhibit excellent performance in drug controlled release, and effectively suppress the corrosion of the substrate.

Means for Solving the Problems

[0006] The technical solution of the present invention provides a coating that has very low requirements for any of the coating manufacturing process, the solvent used in the coating manufacturing process, and the crystal form and size of the drug contained in the final coating, and has excellent performance in drug controlled release through a special design of the coating.

[0007] In the above technical solution of the present invention, the coating includes a drug loading layer and a drug controlled release layer coated outside the drug loading layer, and both the drug loading layer and the drug controlled release layer contain a polymer. The molecular weight retention rate m of the polymer in the drug controlled release layer, the weight average molecular weight M W制御放出 of the polymer, and the thickness X 制御放出 of the polymer satisfy the following relational expression: JPEG2025524216000002.jpg22146 Here, A1 is a fitting constant and is 823. A2 is a fitting constant and is 79.9. k is a fitting constant and is 1.5. b is a fitting constant in the range of 0 ≦ b ≦ 2. The M W制御放出 is the weight average molecular weight of the polymer in the drug controlled release layer, with the unit of kilodalton, and the unit of X 制御放出 is μm.

[0008] In addition, b in the formula of the above technical solution may be any value within the range of 0 to 2. For example, in some embodiments of the present invention, b may be any value among 0.2, 0.5, 0.55, 0.6, 0.65, 0.7, 0.8, 0.9, 1.0, 1.1, 1.2, 1.3, 1.4, 1.5, 1.6, 1.7, 1.8, 1.85, 1.9, 1.95, or 2.0. In other embodiments of the present invention, b may be a range consisting of any two values within the range of 0 to 2, for example, 0.2 to 1.9, 0.5 to 1.5, 0.3 to 1.6, 0.3 to 1.2, 0.55 to 1.95, 0.6 to 1.95, 0.6 to 1.9, 0.6 to 1.85, 0.6 to 1.8, 0.6 to 1.7, 0.7 to 1.8, 0.75 to 1.6, 0.8 to 1.95, 0.8 to 1.9, 0.8 to 1.85, 0.8 to 1.8, 0.8 to 1.7, or 0.8 to 1.6, etc.

[0009] Based on a large number of experimental records, the present invention has found an inherent correlation or regularity among the three variables of the outer drug-controlled release layer, namely the coating thickness, the type of polymer, and the molecular weight of the polymer. According to the specific polymer and the molecular weight of the specific polymer in the outer drug-controlled release layer, there exists a corresponding relatively optimal range of polymer thickness. Within this optimal range of polymer thickness, the performance of the coating in drug release can be significantly improved. The present invention innovatively summarizes the relationship between the optimal thickness range of the drug-controlled release layer and the type of polymer and the molecular weight of the polymer through a large number of experimental results. That is, through the above special design of the outer drug-controlled release layer, the present application relaxes the requirements for the process related to the drug in the drug coating, the solvent used, and the final drug crystal form, and enables the medical device to meet the requirements corresponding to both the drug release rate and the curve basically.

[0010] The coating for the medical device according to the present invention is a relatively general-purpose drug coating and is suitable for both the controlled release of drugs in degradable medical devices and non-degradable medical devices. Preferably, the above coating is suitable for degradable medical devices.

[0011] According to the above technical solution of the present invention, the weight average molecular weight M of the polymer in the drug-controlled release layer W制御放出 is 20 to 900 kDa. In the technical solution of the present invention, the molecular weight of the polymer in the drug-controlled release layer may be any value from 20 to 900 kDa. In some embodiments, the molecular weight of the polymer in the drug-controlled release layer is 21 kDa. In some other embodiments, the molecular weight of the polymer in the drug-controlled release layer is 898 kDa. In some other embodiments, the molecular weight of the polymer in the drug-controlled release layer is 800 kDa. In still some other embodiments, the molecular weight of the polymer in the drug-controlled release layer is 700 kDa.

[0012] According to the above technical solution of the present invention, the molecular weight retention rate m of the polymer in the drug-controlled release layer is 50% to 80%. The molecular weight retention rate m of the polymer described in the present invention refers to the retention rate of the molecular weight of the polymer when the polymer coating is left in a PBS solution (pH = 7.4) at 37 °C for one month. That is, the molecular weight retention rate m = M W制御放出1か月 / M W制御放出 × 100%, where M W制御放出1か月 refers to the molecular weight of the polymer when the polymer coating is left in a PBS solution (pH = 7.4) at 37 °C for one month, and M W制御放出 is the molecular weight of the polymer in the drug-controlled release layer. Each polymer shows a different molecular weight retention rate at the same time point under the same conditions. That is, in the present application, each molecular weight retention rate represents a specific polymer, and different molecular weight retention rates represent different polymers.

[0013] The drug-carrying layer in the present invention may be a single layer or multiple layers, and the drugs carried in the same drug-carrying layer may be one type or multiple types. For example, in some embodiments, the drug-carrying layer is a single layer and the drug carried is one type; in some other embodiments, the drug-carrying layer is three layers, and each layer carries a different drug; in some other embodiments, the drug-carrying layer is two layers, one layer carries one type of drug, and the other layer carries two different types of drugs. That is, a total of three types of drugs are carried in the drug-carrying layer. Further, in some embodiments, the drug-carrying layer is two layers, one layer carries two types of drugs, and the other layer carries two different types of drugs. The two types of drugs carried in each of the two layers may be completely different, partially different, or exactly the same.

[0014] According to the above technical solution of the present invention, the mass ratio of the drug to the polymer in the drug-carrying layer is 1:0.1 to 50. Further, the mass ratio of the drug to the polymer in the drug-carrying layer is 1:(0.1 to 30) or 1:0.1 to 20. Further, the mass ratio of the drug to the polymer in the drug-carrying layer is 1:(0.2 to 18), 1:(0.5 to 18), 1:(0.6 to 15), 1:(0.8 to 18), 1:(0.8 to 16), 1:(0.8 to 15), 1:(0.5 to 12), 1:(0.5 to 8), 1:(0.5 to 6), 1:(0.2 to 16), 1:(0.2 to 12), 1:(0.1 to 12), 1:(0.2 to 10), 1:(0.1 to 10), 1:(0.2 to 8), 1:(0.1 to 8), 1:(0.2 to 5), 1:(0.2 to 4), 1:(0.2 to 3), 1:(0.1 to 5), 1:(0.1 to 4), or 1:(0.5 to 4).

[0015] According to the above technical solution of the present invention, the particle size of the drug particles in the drug loading layer is r≤0.8 μm. The crystal form of the drug in the drug loading layer may be crystalline, amorphous, or semi-crystalline. That is, in some embodiments of the present invention, the drug in the drug loading layer is crystalline, in some other embodiments, the drug in the drug loading layer is amorphous, and in still some other embodiments, the drug in the drug loading layer is semi-crystalline. Due to the special outer layer coating design of the present invention, in the present invention, there are no high requirements regarding the crystal form and crystal grain size of the drug in the drug loading layer. That is, regardless of whether the drug in the drug loading layer is in a crystalline form, an amorphous form, or a semi-crystalline form, as long as the particle size of the drug is controlled to be 0.8 μm or less, the drug release rate can meet the corresponding needs. Therefore, in the present application, due to the special design of the outer layer drug controlled release layer and the drug loading layer, the requirements for the drug loading process and the solvent for the drug loading layer are relatively significantly reduced. As a result, in the case of a drug coating formed through various processes and solvents, as long as the particle size of the drug is not too large and the dispersion is relatively uniform, the drug release rate will be within a relatively ideal range.

[0016] In some embodiments of the present invention, the particle size of the drug particles in the drug loading layer is r≤0.8 μm, in some other embodiments, the particle size of the drug particles in the drug loading layer is r≤0.6 μm, and in some other embodiments of the present invention, the particle size of the drug particles in the drug loading layer is r≤0.45 μm.

[0017] It should be noted that in the present invention, the particle size means the average particle size. For example, "the particle size r of the drug particles in the drug loading layer is 0.6 μm" means that the average particle size of the drug particles in the drug loading layer is 0.6 μm, and there are drug particles with a particle size larger than 0.6 μm and also drug particles with a particle size smaller than 0.6 μm.

[0018] It should be noted that when the drug in the drug loading layer is in a crystalline form, the particle size of the drug particles is the particle size of the crystalline form. When the drug in the drug loading layer is in an amorphous or semi-crystalline form, the particle size of the drug particles is the particle size of the drug clusters dispersed in the polymer.

[0019] According to the above technical solution of the present invention, the total thickness of the drug loading layer and the drug controlled release layer is 2.5 μm or more. Further, the total thickness of the drug loading layer and the drug controlled release layer is 5 μm or more. Further, the total thickness of the drug loading layer and the drug controlled release layer is between (5 μm, 38 μm]. In some embodiments of the present invention, the total thickness of the drug loading layer and the drug controlled release layer is 5.1 μm, 5.5 μm, 5.2 μm, or 6 μm. In some embodiments of the present invention, the total thickness of the drug loading layer and the drug controlled release layer is 7 μm, 7.5 μm, or 8 μm.

[0020] According to the above technical solution of the present invention, the molecular weight of the polymer in the drug loading layer is 30 to 1000 kDa. Further, the molecular weights of the polymers in the inner layer and the intermediate layer are 40 to 1000 kDa, 50 to 1000 kDa, 60 to 1000 kDa. Further, the molecular weight of the polymer in the drug loading layer is 70 to 1000 kDa, 80 to 1000 kDa.

[0021] According to the above technical solution of the present invention, the drug loading layer includes a corrosion control layer on the inner side. The molecular weight of the polymer in the corrosion control layer is 50 to 1000 kDa. Further, the molecular weights of the polymers in the inner layer and the intermediate layer are 60 to 1000 kDa, 70 to 1000 kDa, or 80 to 1000 kDa. Further, the molecular weight of the polymer in the drug loading layer is 90 to 1000 kDa, 100 to 1000 kDa.

[0022] In the above technical solution according to the present invention, the molecular weight of the polymer in the drug-controlled release layer may vary within a wide range. The drug-carrying layer and the corrosion control layer have higher requirements for the molecular weight of the polymer than the drug-controlled release layer. The polymer in the drug-carrying layer is a carrier of the active drug, and its type, molecular weight, thickness of the drug-carrying coating, and mass ratio to the drug all have an important influence on the drug release rate. The higher the molecular weight of the polymer, the clearer the phase separation between the polymer and the drug, the faster the drug release rate, and the smaller the ratio of the mass of the polymer to the mass of the drug, the weaker the ability of the polymer to bind the drug, and the faster the drug release. The thinner the drug-carrying coating, the shorter the diffusion release path of the drug, and the faster the drug release rate. In the corrosion control layer, the molecular weight of the polymer is relatively high, but if the molecular weight is too small, it will corrode rapidly in the initial stage, and the degradable metal part of the device is likely to corrode rapidly in the initial stage, unable to meet the requirements of mechanical properties, or the polymer is decomposed early, and then the rapid corrosion of the degradable metal part of the device may not be effectively controlled.

[0023] According to the above technical solution of the present invention, the total thickness of the corrosion control layer, the drug-carrying layer and the drug-controlled release layer is 3.5 - 45 μm. Further, the total thickness of the corrosion control layer, the drug-carrying layer and the drug-controlled release layer is 5.5 - 40 μm. Further, the total thickness of the corrosion control layer, the drug-carrying layer and the drug-controlled release layer is 5.8 - 36 μm.

[0024] According to the above technical solution of the present invention, the thickness ratio of the corrosion control layer, the drug-carrying layer and the drug-controlled release layer is 1:(0.5 - 15):(0.2 - 13).

[0025] According to the above technical solution of the present invention, the thickness of the drug-controlled release layer is 0.2 - 8 μm. Further, the thickness of the corrosion control layer is 0.5 - 7 μm. Further, the thickness of the corrosion control layer is 0.8 - 6.6 μm.

[0026] According to the above technical solution of the present invention, the thickness of the corrosion control layer is 0.2 to 6 μm, further, the thickness of the corrosion control layer is 0.5 to 5 μm, and further, the thickness of the corrosion control layer is 0.8 to 3 μm.

[0027] According to the above technical solution of the present invention, the thickness of the drug loading layer is 1 to 8 μm, further, the thickness of the corrosion control layer is 1.5 to 7 μm, and further, the thickness of the corrosion control layer is 1.5 to 6 μm.

[0028] According to the above technical solution of the present invention, in the three-layer structure of the coating, at least one layer also contains a corrosion accelerator, a corrosion inhibitor, or an antioxidant.

[0029] According to the above technical solution of the present invention, the coating further includes a metal isolation layer located inside the corrosion control layer. The metal isolation layer is a coating containing a metal element more active than the medical device substrate, and the metal coating can be densely and uniformly distributed on the surface of the degradable device substrate by methods such as electroplating, spray plating, or electroless plating.

[0030] According to the above technical solution of the present invention, the metal isolation layer can slow down the corrosion of the medical device substrate.

[0031] In some embodiments of the present invention, the electronegativity of at least one metal in the metal isolation layer is lower than that of the metal of the substrate of the degradable medical device, that is, at least one metal in the metal isolation layer is more reactive than the metal of the substrate of the degradable medical device. The metal isolation layer may be a pure metal or a metal alloy. When the metal isolation layer is a pure metal, the electronegativity of this metal may be smaller than that of the metal of the substrate of the degradable medical device, that is, this metal is more reactive than the metal of the substrate of the degradable medical device. For example, when the substrate of the degradable medical device is made of iron-based material, the metal isolation layer may be pure zinc or pure magnesium. When the metal isolation layer is a metal alloy, the main metal element in the metal alloy, that is, one or more elements having a higher content, is more reactive or has a lower electronegativity than the metal of the substrate of the degradable medical device. For example, when the substrate of the degradable medical device is made of iron-based material, the metal isolation layer may be a zinc alloy or a magnesium alloy. In some other embodiments of the present invention, the metal isolation layer may be an isolation layer containing metal oxides. In some other embodiments of the present invention, the metal isolation layer may be another metal-containing coating that can prevent the corrosion of the medical device and the local environment.

[0032] According to the above technical solution of the present invention, when the metal in the absorbent metal substrate is pure iron or an iron-based alloy, the metal isolation layer includes pure magnesium, a magnesium-containing alloy, pure zinc or a zinc-containing alloy. When the metal in the absorbent metal substrate is pure zinc or a zinc-containing alloy, the metal isolation layer includes pure magnesium or a magnesium-containing alloy.

[0033] Since the metal in the metal isolation layer is more active than the metal constituting the device substrate, in the microenvironment in vivo, electrochemical protection is formed. The relatively highly active metal in the isolation layer is preferentially corroded, and the device substrate that plays a supporting role is protected and can have its corrosion retarded. For example, when the device substrate is iron or an iron alloy, when a metal coating made of zinc or a zinc alloy is coated on the surface of the device, a galvanic cell is formed in the microenvironment in vivo. The metal coating made of iron or a zinc alloy functions as the negative electrode, loses electrons to become zinc ions, and is preferentially corroded. On the other hand, the device substrate made of iron or an iron alloy functions as the positive electrode and is temporarily protected from corrosion.

[0034] Through the four schemes of the drug-controlled release layer and the drug-carrying layer, or the drug-controlled release layer, the drug-carrying layer and the corrosion control layer, or the drug-controlled release layer, the drug-carrying layer and the metal isolation layer, or the drug-controlled release layer, the drug-carrying layer, the corrosion control layer and the metal isolation layer in the present invention, both the control of drug release and the suppression of substrate corrosion can be achieved. In each scheme, in order to achieve the object of the present invention, the parameters of each layer of the coating can be adjusted or combined with each other. For example, in some embodiments of the present invention, the coating for the medical device is, from the outside to the inside, a drug-controlled release layer and a drug-carrying layer, or a drug-controlled release layer, a drug-carrying layer and a corrosion layer. By adjusting the type of polymer, the molecular weight of the polymer, the thickness of the coating, and the drug-carrying amount of each layer of the above coating in the corresponding region, the final coating has both excellent drug release performance and excellent corrosion suppression performance. After the device is implanted at the corresponding site, no growth or inflammation occurs, and at the same time, no corrosion occurs or only slow corrosion occurs in the first 1 to 6 months, but rapid corrosion occurs after 6 months.

[0035] According to the above technical solution of the present invention, the thickness of the metal isolation layer is 0.6 μm or more, and further, the thickness of the metal isolation layer is 0.6 to 3.8 μm. When the thickness of the metal isolation layer is sufficient, it can effectively prevent the premature corrosion of the metal of the substrate of the degradable medical device, thereby ensuring that the substrate does not corrode within the first 1 to 6 months. Also, the thickness of the metal isolation layer should not be too thick. If it is too thick, it will not be completely corroded within 2 years. Finally, the degradable medical device will not corrode rapidly after 6 months and will not be completely corroded within 2 years.

[0036] According to the above technical solution of the present invention, the polymer in the coating may be at least one of a degradable polyester or a degradable polyanhydride. Further, the polymer described in the present invention is polylactide, poly-L-lactic acid, poly-D-lactic acid, polyhydroxyethyl ester, polyurethane, polyamino acid, poly(lactic acid diglycolic acid), poly-D,L-lactide, polypropylene glycol, polyglycolic acid, polylactic acid glycolic acid, polysalicylic anhydride ester, polytrimethylene carbonate, polycaprolactone, polycaprolactone, polyhydroxyalkanoate, polyacrylate, polysuccinate, poly(β-hydroxybutyric acid) and polyethylene adipate, poly 1,3-bis(p-carboxyphenoxy)propane-sebacic acid, poelcic acid dimer-sebacic acid, and at least one selected from polyfumaric acid-sebacic acid. In some embodiments of the present invention, the polymer in the coating may be a specific one of the above polymers, and the polymers in the drug-controlled release layer, the corrosion control layer, and the corrosion control layer in the coating are all polylactide. In some other embodiments of the present invention, the polymer in the coating may be a mixture of two or more of the above polymers, such as a mixture of two or more of poly-L-lactic acid, poly-D-lactic acid, polyurethane, polysalicylic anhydride ester, polytrimethylene carbonate or polycaprolactone. The mixture of two or more polymers may contain two polymers simultaneously in one layer, or each single layer may contain only one different polymer. In some embodiments of the present invention, the polymer in the outer drug-controlled release layer in the coating is poly-L-lactic acid, the polymer in the drug-carrying layer is polylactide, and the polymer in the corrosion control layer is polyurethane. In some other embodiments of the present invention, the polymers in the drug-controlled release layer, the corrosion control layer, and the corrosion control layer of the coating are all a mixture of polyurethane and polysalicylic anhydride ester, or a mixture of polylactide, poly-L-lactic acid, and poly-D-lactic acid.In some further embodiments, the polymers in the drug-controlled release layer and the corrosion-controlled layer in the coating are poly-DL-lactic acid, and the polymer in the corrosion-controlled layer is a mixture of polyhydroxyalkanoate and polyacrylate.

[0037] According to the above technical solution of the present invention, the drug carried on the coating is at least one of an angiogenesis inhibitor, an anti-inflammatory drug, an antithrombotic drug, and a sensitizing agent. The antithrombotic drug includes an antiplatelet drug and an anticoagulant drug. In some embodiments of the present invention, the drug carried on the coating is only an angiogenesis inhibitor. In some other embodiments, the drug carried on the coating is a combination of any two or more drugs among an angiogenesis inhibitor, an anti-inflammatory drug, an antithrombotic drug, and a sensitizing agent. More specifically, the angiogenesis inhibitor described in the present invention is at least one selected from paclitaxel, rapamycin, and their derivatives, the antiplatelet drug includes cilostazol, etc., the antithrombotic drug includes heparin, the anti-inflammatory drug is dexamethasone, and the anti-allergy drug is at least one selected from calcium gluconate, chlorpheniramine, and cortisone.

[0038] According to the above technical solution of the present invention, the drug carried on the drug-carrying layer is at least one of paclitaxel, rapamycin, cilostazol, heparin, dexamethasone, tacrolimus, everolimus, calcium gluconate, chlorpheniramine, or cortisone.

[0039] According to the above technical solution of the present invention, the medical device is a degradable metal medical device or a degradable non-metal medical device, and the metal is a pure metal or a metal alloy. Further, the degradable medical device is a medical device that at least partially includes iron-based, zinc-based, polymer, and magnesium-based. In some embodiments of the present invention, the substrate of the degradable medical device is composed of one kind of metal or alloy. For example, the entire degradable medical device is an iron-based medical device. In some other embodiments of the present invention, the substrate of the degradable medical device is composed of two or more materials such that a part thereof is iron-based and the rest is zinc-based, or a part thereof is metal-based and the rest is non-metal-based, etc.

[0040] According to the above technical solution of the present invention, the medical device includes any one of a vascular stent, a heart valve, a non-vascular endovascular stent, an occluder, an orthopedic implant, a dental implant, a respiratory implant, a gynecological implant, an andrological implant, a suture, or a bolt.

[0041] According to the above technical solution of the present invention, the stent described in the present invention is a vascular stent or a stent for a heart valve. Further, the stent described in the present invention is a balloon-expandable stent or a self-expandable stent. Further, the stent described in the present invention includes a coronary stent, a stent below the knee, a superficial femoral stent, an intracranial stent, a carotid stent, a pulmonary artery stent, or a renal artery stent.

[0042] According to the above technical solution of the present invention, the base material of the degradable medical device is pure iron or an iron-based alloy with a carbon content of 2.11 wt% or less.

[0043] Regarding the above iron-based, zinc-based, or magnesium-based, the metal therein refers to the main component of the substrate. For example, the iron-based includes pure iron or an iron alloy, that is, the iron-based means that the main component of such a substrate is iron.

[0044] The present invention controls the release of the drug in the drug-carrying layer through the relationship or internal logic among the type of polymer, the molecular weight of the polymer, and the thickness of the polymer in the drug-controlled release layer. When the stent is first implanted into the human body, the drug in the drug-carrying layer is rapidly released by the erosion of blood, and the initial drug release is too fast, resulting in insufficient drug release in the later stage and the inability to release the drug effectively and continuously. The drug-controlled release layer of the present invention effectively avoids this situation, thereby avoiding restenosis of blood vessels. That is, the drug-controlled release layer of the present invention not only has a certain protective effect on the drug-carrying layer but also can adjust the release rate of the drug carried by the drug-carrying layer. The present invention adds a special drug-controlled release layer outside the drug-carrying layer. Only when the polymer in the drug-controlled release layer is decomposed to a certain extent, the drug in the drug-carrying layer gradually increases the contact area between the drug-carrying layer and the blood, so that the problem of burst release of the drug in the initial stage of implanting the stent can be effectively suppressed.

[0045] The drug-controlled release layer and the drug-carrying layer in the present invention cooperate to control the release of the drug, thereby effectively controlling the release rate of the drug at each stage and ensuring that the amount of drug released at each stage meets the demand. On the other hand, in order to prevent excessive drug loading from causing any toxicity or side effects to the human body, the drug-carrying layer should carry an appropriate amount of drug as much as possible. That is, while controlling the release of the drug so that the drug amount meets the demand, it is necessary to reduce the total drug loading amount as much as possible.

[0046] The drug-controlled release layer, the drug-carrying layer, and the corrosion-promoting layer in the present invention all contain polymers. Through the synergistic effect of the thickness and molecular weight of the polymers in the three coatings, the corrosion of the device substrate, or the device substrate and the metal isolation layer, is controlled. In some embodiments according to the present invention, when the molecular weight of the polymer in the drug-controlled release layer is relatively small, in order to control the final comprehensive corrosion rate of the device substrate, the types and molecular weights of the polymers in the drug-carrying layer and the corrosion control layer can be adjusted to meet the requirements. For example, increase the molecular weight of the polymers in the drug-carrying layer and the corrosion control layer and select polymers with a relatively slow decomposition rate.

[0047] The metal isolation layer introduced into the technical solution of the present invention, together with the above-mentioned polymer-containing coating, can control the corrosion rate of the device substrate to ensure that the device is not corroded in the initial stage of implantation (1 to 6 months). When the blood vessel is repaired, the substrate of the device is rapidly decomposed. In the in-vivo environment, when the corrosion control layer and the metal isolation layer are decomposed simultaneously and the degradable polyester constituting the corrosion control layer is decomposed, acidic substances are generated. When this decomposition accumulates to a certain extent, a local acidic environment is generated, which may accelerate the corrosion and decomposition of the metal isolation layer and the substrate of the medical device. This process can be controlled by the thickness and molecular weight of the degradable polyester of the corrosion control layer. The lower the molecular weight, the thicker the degradable polyester, and the faster the corrosion and decomposition of the metal isolation layer and the stent.

[0048] The polymers in the corrosion control layer, drug carrier layer, and drug controlled release layer in the present invention cooperate to control the corrosion of the device substrate or the device substrate and the metal isolation layer. In addition, the corrosion control layer also serves to isolate metal ions generated by the device substrate and the metal isolation layer from the human body environment during the corrosion process, thereby preventing metal ions combined with the active drug from moving to human blood vessels without being inhibited and causing hemolysis, or preventing excessive metal ions from directly contacting the human body and causing local poisoning or thrombosis.

[0049] It should be understood that the terms used in this specification are for the sole purpose of describing specific exemplary embodiments and are not intended to be limiting. As used in this specification, the singular forms "a", "one", and "the" shall include the plural forms as well, unless the context clearly dictates otherwise. The terms "comprising", "including", "containing", and "having" are inclusive and indicate the presence of the recited features, steps, operations, elements, and / or components, but do not preclude the presence or addition of one or more other features, steps, operations, elements, components, and / or combinations thereof. Method steps, processes, and operations described herein should not be construed as necessarily being executed in the particular order described or illustrated, unless the execution order is explicitly indicated. It should also be understood that additional or alternative steps may be used.

Brief Description of the Drawings

[0050] Various other advantages and beneficial effects will become apparent to those skilled in the art upon reading the following detailed description of the preferred embodiments. The drawings are for the sole purpose of illustrating the preferred embodiments and should not be construed as limiting the present invention. Also, throughout the figures, the same parts are denoted by the same reference numerals.

Figure 1

Figure 2

Figure 3

Figure 4

Figure 5

Figure 6

Embodiments for Carrying Out the Invention

[0051] The following description is only a preferred embodiment of the present invention, and the protection of the present invention is not limited to the following preferred embodiments. Although a stent is used as an example in the examples, it does not mean that the technical solution of the present invention is only applicable to stents. For those skilled in the art, any modifications and improvements made based on the concept of this invention are all included within the protection scope of the present invention. When the manufacturers of the reagents and instruments used are not described, all are conventional products commercially available.

[0052] Test Method 1. Measurement of the weight average molecular weight of the polymer For detection, a GPC multi-angle laser light scattering device from Wyatt, USA, was used in combination with a molecular weight measurement system. This measurement system includes a liquid phase pump and a sampler from Agilent, USA, an Agilent PL MIXED-C type GPC column (size: 7.5×300 mm, 5 microns) from Agilent, USA, a multi-angle laser light scattering device and a differential detector from Wyatt, USA. The detection conditions are as follows.

[0053] Mobile phase: Tetrahydrofuran, pump flow rate: 1 mL / min, sample volume: 100 μL, laser wavelength: 663.9 nm, test temperature: 35 °C 2. Measurement of the thickness of the metal isolation layer To measure the thickness of the metal isolation layer, the fluorescent X-ray film thickness measurement method is used. First, it is necessary to calibrate the equipment using the standard block of the corresponding element. After calibration is completed, fix the stent sample for testing the thickness of the metal isolation layer on the sample stage, place it in the fluorescent X-ray film thickness meter, set the types of coating and base metal, set parameters such as the measurement time of 10 - 15 s and the thickness unit of μm. Click [OK] to test the thickness of the metal isolation layer.

[0054] 3. Measurement of the total thickness of the polymer coating For measurement, the scanning electron microscopy method is used. First, fix the stent sample that needs to test the coating thickness on the sample stage. Next, place the sample stage in the JFC-1600 sputter coater and spray platinum. After spraying once, rotate it 180° and spray again so that all positions are sprayed. Put the sample with platinum sprayed on the surface into the reagent mixture of room temperature resin hardener manufactured by Buehler at a ratio of 5:1, leave it for more than 8 hours, and then separate it from the sealing case. Divide the sealed sample into three equal parts, and polish each one according to the polishing procedure of the sample using a semi-automatic polishing machine until the abrasion marks of the sample to be measured disappear. Fix the polished sample on the stage of the scanning electron microscope, place the entire stage in the JFC-1600 sputter coater, and spray for 20 seconds. Put the sprayed sample into the JSM-6510 scanning electron microscope to measure the thickness.

[0055] 4. Test method for the thickness of each layer of the multilayer drug coating First, fix the stent sample that needs to evaluate the characteristics of the multilayer drug coating on the sample stage. Next, place the sample stage into the JFC-1600 sputter coater and spray platinum. After spraying once, rotate it 180° and spray again so that all positions are sprayed. Put the sample with platinum sprayed on its surface into the reagent mixture of room temperature resin hardener manufactured by Buehler at a ratio of 5:1, leave it for more than 8 hours, and then separate it from the sealing case. Divide the sealed sample into three equal parts, and polish one by one according to the polishing procedure of the sample using a semi-automatic polishing machine until the wear marks of the sample to be measured disappear. Fix the polished sample on the stage of a scanning electron microscope, put the whole stage into the JFC-1600 sputter coater, and spray for 20 seconds.

[0056] Use the microscopic Raman spectrometer Thermo DXR2 to scan the cross-section of the coating part of the medical device with the multilayer coating. The laser energy parameter is 6.0 mW, the exposure time is 20 Hz, the number of exposures is 50 times, and the magnification is 500 times.

[0057] By YZ cross-section scanning, a component distribution diagram of the cross-section can be obtained. From the characteristic peaks of the drug and polylactic acid, the drug and polylactic acid content in different regions can be qualitatively analyzed, and the thickness of each coating can be tested.

[0058] 5. Test method for the particle size of drug particles during coating After vacuum drying the coatings composed of different solvents at room temperature for 24 hours, put them into those containing 5 mL of isopropyl alcohol solution, and rotate and shake them at 37 °C and a speed of 100 rpm for 30 minutes. After completely extracting the drug with isopropyl alcohol, take out the coating, vacuum dry it for 24 hours, and put it into liquid nitrogen for brittle fracture treatment. Fix the sample subjected to the brittle fracture treatment on the stage of a scanning electron microscope, put the whole stage into the JFC-1600 sputter coater, and spray for 20 seconds. Set the sprayed sample on a JSM-6510 scanning electron microscope, and measure the particle size of the drug particles by measuring the size of the pores remaining after drug leaching.

[0059] 6. Test method for crystallinity of drug during coating Coating solutions composed of different solvents (ethyl acetate, chloroform) are uniformly sprayed onto a stainless-steel plate using a spraying device and vacuum-dried at room temperature for 24 hours to produce films of drug / polymer coatings composed of different solvents.

[0060] XRD tests (Bruker D2 PHASER) are performed on drug / polymer films made of different solvents, and the samples are scanned using the θ-2θ linkage mode. The scanning angle range is 3 - 40°, the scanning step is 0.02°, and the scanning speed is 4° / min.

[0061] 7. Test method for drug release rate of medical devices with multilayer coatings in animals' bodies Medical devices with multilayer coatings are implanted into the blood vessels of healthy rabbits. The rabbits are sacrificed at predetermined observation time points such as 7 hours, 1 day, 3 days, 7 days, 14 days, 28 days, 60 days, 90 days, and 180 days, and the medical devices with multilayer coatings are removed. After removing as much tissue as possible from the medical devices, they are placed in brown glass bottles, an appropriate amount of acetonitrile is added to make a fixed volume, and the volume can be determined according to specific specifications so that the stent can be completely immersed. Ultrasonic treatment is performed for 15 minutes, filtered through a 0.22μm nylon organic filter, placed in the injection bottle of the chromatograph, and the amount of drug remaining on the stent is measured using a liquid chromatograph. The drug release amount of the stent is the difference between the initial drug amount on the stent and the residual drug amount (test drug amount) on the stent after immersion, and the drug release percentage is the ratio of the drug release amount to the initial drug amount.

[0062] Chromatography conditions: Mobile phase: acetonitrile: water = 65:35 solution, flow rate 1.0 ml / min, detection wavelength 278 nm, column temperature 50°C, injection volume 10.0 μl.

[0063] 8. Test for mass loss of absorbable stents in animals' bodies A medical device with a multi-layer coating was implanted into the blood vessels of healthy rabbits. The rabbits were sacrificed at scheduled observation time points such as 7 hours, 1 day, 3 days, 7 days, 14 days, 28 days, 60 days, 90 days, and 180 days later, and the medical device with the multi-layer coating was removed. After removing as much tissue as possible from the medical device, a liquid chromatograph was used to test the amount of drug remaining on the stent. The stent on which the amount of drug was tested was washed with an ethyl acetate solution to completely dissolve the polymer coating on the stent.

[0064] Perform micro-CT examination on the stent washed above and conduct qualitative analysis of the corrosion of the iron-based substrate. Then, place the stent in a tartaric acid solution, perform ultrasonic cleaning to remove the surface corrosion product layer, and after drying, weigh it as m 腐食後 Let it be.

[0065] The mass loss rate m% of the device is the ratio of the difference between the mass m 腐食前 of the device before corrosion and the mass m 腐食後 of the device after corrosion and treatment to the mass m 腐食前 of the device before corrosion, multiplied by 100%. That is, it is as follows. JPEG2025524216000003.jpg12146

[0066] 9. Test of the radial support strength of the stent The stent is expanded to the nominal pressure using an appropriate balloon (or distal component) so that the outer diameter of the stent sample reaches a predetermined size.

[0067] The length of the stent is measured using a three-dimensional microscope, and a test of the support force curve is performed with the 50% compression displacement as the end point. The radial support strength is defined as the value of the force per unit length when the nominal diameter of the stent is compressed by 10% as an expression of the value of the radial extrusion resistance of the stent (unit: kPa).

[0068] <Example 1> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 0.6 μm was plated on the surface of the base material by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 3 μm was uniformly sprayed outside the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 600 kDa was used for the coating, and chloroform was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-carrying layer with an average thickness of 5 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 400 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer was 1:0.5, the average particle diameter of the drug particles was 0.25 μm, and the drug particles were amorphous (see Figure 3). After drying the drug-carrying layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 3 μm was uniformly sprayed. For the coating with a molecular weight retention rate of 74% at one month and a molecular weight of 200 kDa, poly(lactide-co-glycolide) was used, and ethyl acetate was used as the solvent.

[0069] In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not blocked, and there was no restenosis. Due to the coating, the stent obtained sufficient support within 6 months and was completely degraded within 2 years.

[0070] <Example 2> In this example, an iron-based stent with a standard size of 3.0 mm×8 mm was used. A zinc layer with an average thickness of 2 μm was plated on the surface of the substrate as a metal isolation layer by electroplating. Using a stent spraying device, a drug-carrying layer with an average thickness of 10 μm was uniformly sprayed outside the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 400 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer was 1:8, the average particle size of the drug particles was 0.35 μm, and the drug particles were amorphous. After drying the drug-carrying layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 1.8 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 400 kDa and a molecular weight retention rate of 79% in one month was used for the coating, and chloroform was used as the solvent.

[0071] In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not blocked, and there was no restenosis. Due to the coating, the stent obtained sufficient support within 6 months and was completely degraded within 2 years.

[0072] <Example 3> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 3 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 6 μm was uniformly sprayed outside the metal isolation layer. PLGA (50:50) with a molecular weight of 50 kDa was used for the coating, and chloroform was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-loading layer with an average thickness of 5 μm was uniformly sprayed. PLGA (50:50) with a molecular weight of 50 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be loaded. The mass ratio of sirolimus to PLGA in the drug-loading layer was 1:1.25, the average particle diameter of the drug particles was 0.45 μm, and the drug particles were amorphous. After drying the drug-loading layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 6 μm was uniformly sprayed. PLGA (50:50) with a molecular weight of 50 kDa and a molecular weight retention rate of 61% after one month was used for the coating, and chloroform was used as the solvent.

[0073] In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not blocked, and there was no restenosis. Due to the coating, the stent obtained sufficient support within 6 months and was completely degraded within 2 years.

[0074] <Example 4> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 3.8 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 5 μm was uniformly sprayed outside the metal isolation layer. Poly (lactide-co-glycolide) with a molecular weight of 400 kDa was used for the coating, and ethyl acetate was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-carrying layer with an average thickness of 4 μm was uniformly sprayed. Poly (lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, ethyl acetate was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly (lactide-co-glycolide) in the drug-carrying layer is 1:2, the average particle diameter of the drug particles is 0.2 μm, and the drug particles are semi-crystalline. After drying the drug-carrying layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 0.8 μm was uniformly sprayed. Poly (lactide-co-glycolide) with a molecular weight of 800 kDa and a molecular weight retention rate of 81% in one month was used for the coating, and chloroform was used as the solvent.

[0075] In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not blocked, and there was no restenosis. Due to the coating, the stent obtained sufficient support within 6 months and was completely degraded within 2 years.

[0076] <Example 5> In this example, an iron-based stent with a standard size of 3.0 mm×8 mm was used. A zinc layer with an average thickness of 1 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spray device, a corrosion control layer with an average thickness of 3.5 μm was uniformly sprayed outside the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, and ethyl acetate was used as the solvent. After drying the corrosion control layer, using a stent spray device, a drug-loading layer with an average thickness of 5 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 100 kDa was used for the coating, ethyl acetate was used as the solvent, and sirolimus was used as the drug to be loaded. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-loading layer was 1:1.5, the average particle diameter of the drug particles was 0.7 μm, and the drug particles were semi-crystalline. After drying the drug-loading layer, using a stent spray device, a drug-controlled release layer with an average thickness of 5.8 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 50 kDa and a molecular weight retention rate of 63% after 1 month was used for the coating, and ethyl acetate was used as the solvent.

[0077] In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not occluded, and there was no restenosis. Due to the coating, the stent obtained sufficient support within 6 months and was completely degraded within 2 years.

[0078] <Example 6> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 2.5 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 2.5 μm was uniformly sprayed outside the metal isolation layer. Poly (rac-lactic acid) with a molecular weight of 300 kDa was used for the coating, and ethyl acetate was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-loading layer with an average thickness of 8 μm was uniformly sprayed. Poly (rac-lactic acid) with a molecular weight of 300 kDa was used for the coating, ethyl acetate was used as the solvent, and sirolimus was used as the drug to be loaded. The mass ratio of sirolimus to poly (rac-lactic acid) in the drug-loading layer is 1:6, the average particle diameter of the drug particles is 0.35 μm, and the drug particles are semi-crystalline (see Figure 3). After drying the drug-loading layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 2.6 μm was uniformly sprayed. Poly (rac-lactic acid) with a molecular weight of 600 kDa, a molecular weight retention rate of 80% after 1 month, and chloroform was used as the solvent. In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not blocked, and there was no restenosis.

[0079] <Example 7> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 1.5 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 2 μm was uniformly sprayed outside the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, and acetone was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-carrying layer with an average thickness of 3 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 600 kDa was used for the coating, acetone was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer is 1:1.5, and the average particle diameter of the drug particles is 0.6 μm. After drying the drug-carrying layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 6.6 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 100 kDa, a molecular weight retention rate of 65% after 1 month, and acetone was used as the solvent.

[0080] In this example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Throughout the implantation stage, no thrombus was observed in the blood vessel where the stent was placed, the lumen was not blocked, and there was no restenosis. Due to the coating, the stent obtained sufficient support within 6 months and was completely degraded within 2 years.

[0081] <Comparative Example 1> In this example, an iron-based stent with a standard size of 3.0 mm×8 mm was used. A zinc layer with an average thickness of 1 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 3 μm was uniformly sprayed outside the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 400 kDa was used for the coating, and ethyl acetate was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-carrying layer with an average thickness of 5 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, ethyl acetate was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer is 1:2, the average particle size of the drug particles is 0.3 μm, and the drug particles are semi-crystalline.

[0082] In this comparative example, the drug release rate of the stent in vivo is shown in Fig. 4. As can be seen from Fig. 4, in this comparative example, burst release occurred in the stent at the initial stage, and most of the drug was released at the initial stage of stent implantation. Therefore, the drug release became insufficient in the middle and late stages. On the 90th day, serious signs of proliferation were observed in the blood vessel where the stent was placed. The results of the mass loss of the stent are shown in Fig. 5, and the radial support strength of the stent is shown in Fig. 6. In vivo, the stent corrodes at a normal rate, so it can provide effective support even at the 6-month point.

[0083] <Comparative Example 2> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 1.5 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spray device, a corrosion control layer with an average thickness of 5 μm was uniformly sprayed outside the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, and chloroform was used as the solvent. After drying the corrosion control layer, using a stent spray device, a drug-carrying layer with an average thickness of 5 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer was 1:1.25. Poly(lactide-co-glycolide) with a molecular weight of 400 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be carried. The average particle diameter of the drug particles was 0.35 μm, and the drug particles were amorphous.

[0084] In this comparative example, the drug release rate of the stent in vivo is shown in Fig. 4. As can be seen from Fig. 4, in this comparative example, burst release occurred in the stent at the initial stage, and most of the drug was released at the initial stage of stent implantation. Therefore, the drug release became insufficient in the middle and late stages. On the 90th day, serious signs of proliferation and thrombosis were observed in the blood vessel where the stent was placed.

[0085] Refer to Fig. 5 for the mass loss rate diagram of the stent, and the radial support strength of the stent is shown in Fig. 6. In vivo, the stent corrodes at a normal rate, so it can provide effective support even at the 6-month point.

[0086] <Comparative Example 3> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 1 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spray device, a corrosion control layer with an average thickness of 4 μm was uniformly sprayed outside the metal isolation layer. Polylactide with a molecular weight of 800 kDa was used for the coating, and chloroform was used as the solvent. After drying the corrosion control layer, using a stent spray device, a drug-carrying layer with an average thickness of 4 μm was uniformly sprayed. Polylactide with a molecular weight of 400 kDa was used for the coating, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to polylactide in the drug-carrying layer is 1:2, the average particle diameter of the drug particles is 0.35 μm, and the drug particles are semi-crystalline. Ethyl acetate was used as the solvent. After drying the drug-carrying layer, using a stent spray device, a drug-controlled release layer with an average thickness of 15 μm was uniformly sprayed. Polylactide with a molecular weight of 200 kDa and a molecular weight retention rate of 74% after one month was used for the coating, and ethyl acetate was used as the solvent.

[0087] In this comparative example, the drug release rate of the stent in vivo is shown in Figure 4. As can be seen from Figure 4, in this comparative example, since the drug release of the stent is sufficient in the early stage, it cannot play a therapeutic role in the initial stage of implantation. On the 14th day, signs of thrombosis were observed in the blood vessel where the stent was placed.

[0088] The mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. As the thickness of the coating increases, the amount of polylactic acid increases accordingly, the corrosion of the stent in vivo speeds up, the drug release becomes slower in the initial stage, and after 60 days of implantation, it cannot maintain a normal shape and cannot provide effective support. On the 30th day, obvious thrombus was observed on the stent rod, and after 90 days of implanting the stent, the blood vessel had severe restenosis and the lumen area was lost by more than 60%.

[0089] <Comparative Example 4> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. A zinc layer with an average thickness of 0.6 μm was plated on the surface of the substrate by electroplating as a metal isolation layer. Using a stent spraying device, a corrosion control layer with an average thickness of 2.5 μm was uniformly sprayed on the surface of the metal isolation layer. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, and chloroform was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-carrying layer with an average thickness of 8 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 200 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer is 1:8, the average particle diameter of the drug particles is 0.35 μm, and the drug particles are amorphous. After drying the drug-carrying layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 0.2 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 100 kDa and a molecular weight retention rate of 68% after one month was used for the coating, and ethyl acetate was used as the solvent.

[0090] In this comparative example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. The stent corrodes at a normal rate in vivo and can provide effective support even at the 6-month mark. However, in this comparative example, since the drug-controlled release layer of the stent is too thin, initial burst release of the drug occurs, and most of the drug is released in the early stage of stent implantation. Therefore, the drug release becomes insufficient in the middle and late stages. On the 90th day, serious signs of proliferation were observed in the blood vessel where the stent was placed.

[0091] <Comparative Example 5> In this example, an iron-based stent with a standard size of 3.0 mm × 8 mm was used. Using a stent spraying device, a coating with an average thickness of 2.5 μm was uniformly sprayed onto the surface of the metal substrate. Poly(lactide-co-glycolide) with a molecular weight of 100 kDa was used for the coating, and chloroform was used as the solvent. After drying the corrosion control layer, using a stent spraying device, a drug-carrying layer with an average thickness of 5 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 100 kDa was used for the coating, chloroform was used as the solvent, and sirolimus was used as the drug to be carried. The mass ratio of sirolimus to poly(lactide-co-glycolide) in the drug-carrying layer was 1:0.5, the average particle diameter of the drug particles was 0.25 μm, and the drug particles were amorphous. After drying the drug-carrying layer, using a stent spraying device, a drug-controlled release layer with an average thickness of 0.2 μm was uniformly sprayed. Poly(lactide-co-glycolide) with a molecular weight of 100 kDa and a molecular weight retention rate of 68% after one month was used for the coating, and ethyl acetate was used as the solvent.

[0092] In this comparative example, the drug release rate of the stent in vivo is shown in Figure 4, the mass loss of the stent is shown in Figure 5, and the radial support strength of the stent is shown in Figure 6. Since there is no metal isolation layer, the substrate of the stent corrodes rapidly, and the stent rod breaks at the initial stage of implantation. As a result, the release of the drug also accelerates, and most of the drug is released at the initial stage of implanting the stent. After 60 days of implantation, it cannot maintain a normal form and cannot provide effective support. After 30 days, obvious thrombi occur in the stent rod, the tissue around the stent rod turns abnormal in color, and after 90 days of implanting the stent, the blood vessel is severely restenosed, and the lumen area has lost more than 60%.

[0093] From the above, when comparing Example 4 and Comparative Example 1, in Comparative Example 1, due to the absence of a drug-controlled release layer, drug burst release occurs, a large amount of drug is released in a short time, and as a result, in the middle and late stages, almost no drug is released over a very long period, and the drug utilization rate also decreases.

[0094] Comparative Example 1 has the same mass ratio of drug to polyester as Example 4, but since there is no drug-controlled release layer, a large amount of drug is released in a short time. As a result, in the middle and late stages, almost no drug is released over a very long period, and the drug utilization rate is also low. When compared with Comparative Example 2, in Comparative Example 2, although the shape and type of drug particles are different, since there is no drug-controlled release layer in either case, the drug is burst released in vivo.

[0095] Compared with Example 4, Comparative Example 3 has the same solvent and drug particle morphology, but since the drug-controlled release layer is too thick, drug release in the initial and middle stages is insufficient. As a result, blood vessel proliferation and restenosis occur, and there is no therapeutic effect of the drug. In Comparative Example 3, the polylactic acid coating is too thick, and the content of polylactic acid increases, so the stent corrodes faster in vivo. After 60 days of implantation, the stent cannot maintain its normal form, cannot provide effective support, and the symptoms of blood vessel thrombosis and restenosis worsen.

[0096] When comparing Comparative Example 4 and Comparative Example 5, both have a drug-controlled release layer, but since the thickness of the drug-controlled release layer is insufficient, after the polylactic acid on the surface is decomposed, a large amount of drug is still released in a short time. As a result, in the middle and late stages, almost no drug is released over a very long period, and the drug utilization rate is also low.

[0097] In Comparative Example 5, there is no metal isolation layer, and there are no corresponding polymer types and polymer molecular weights. Therefore, the polymer directly contacts the stent substrate, the stent matrix corrodes rapidly, the stent rod breaks in the initial stage, and as a result, drug release is also accelerated. After more than 60 days of implantation, it cannot maintain its normal form and cannot provide effective support.

[0098] In Examples 1 to 7, the molecular weights are different, the polymer materials are different, the reagents are different, the crystallinities are different, and the drug particle diameters are different. However, after preparing a drug-controlled release layer with an appropriate thickness, early burst release of the drug does not occur, and the drug is appropriately released during the release cycle, which more appropriately matches the degradation cycle of the stent substrate.

[0099] The present invention can also have various other embodiments, and those skilled in the art can make various corresponding changes and modifications according to the present invention without departing from the spirit and essence of the present invention. However, all these corresponding changes and modifications should fall within the scope of the claims appended to the present invention.

Claims

1. A coating for a medical device comprising a drug-carrying layer and a drug-controlled release layer covering the outside of the drug-carrying layer, wherein both the drug-carrying layer and the drug-controlled release layer contain polymers, The molecular weight retention rate m of the polymer in the drug-controlled release layer, and the weight-average molecular weight M of the polymer. Ｗ制御放出 , and polymer thickness X 制御放出 The following relationship is satisfied, Here, A １ This is the fitting constant, which is 823. A ２ This is the fitting constant, which is 79.

9. k is a fitting constant, which is 1.

5. b is a fitting constant in the range 0 ≤ b ≤ 2, Said M Ｗ制御放出 X is the weight-average molecular weight of the polymer in the drug-controlled release layer, with units of kilodaltons. 制御放出 A coating for medical devices characterized by having a unit of μm.

2. The weight-average molecular weight M of the polymer in the drug-controlled release layer. Ｗ制御放出 The coating for medical devices according to claim 1, characterized in that the pressure is 20 to 900 kDa and the molecular weight retention rate m of the polymer is 50% to 80%.

3. The coating for a medical device according to claim 1, characterized in that the mass ratio of drug to polymer in the drug-carrying layer is 1:(0.1 to 50).

4. The coating for a medical device according to claim 1, characterized in that the particle size r of the drug particles in the drug-carrying layer is 0.8 μm or less.

5. The coating for a medical device according to claim 1, characterized in that the total thickness of the drug-carrying layer and the drug-controlled release layer is 2.5 μm or more.

6. The coating for medical devices according to claim 1, characterized in that the molecular weight of the polymer in the drug-carrying layer is 30 to 1000 kDa.

7. The coating for a medical device according to claim 1, characterized in that the inside of the drug-carrying layer further comprises a corrosion control layer, and the molecular weight of the polymer in the corrosion control layer is 50 to 1000 kDa.

8. The coating for medical devices according to claim 7, characterized in that the total thickness of the corrosion control layer, drug carrying layer, and drug control release layer is 3.5 to 45 μm.

9. The coating for medical devices according to claim 7, characterized in that the thickness ratio of the corrosion control layer, the drug carrying layer, and the drug control release layer is 1:(0.5 to 15):(0.2 to 13).

10. The coating for a medical device according to claim 7, further comprising a metal isolation layer located inside the corrosion control layer.

11. The coating for medical devices according to claim 10, characterized in that the thickness of the metal isolation layer is 0.6 μm or more.

12. The polymer is at least one of biodegradable polyesters or biodegradable polyanhydrides. The coating for medical devices according to claim 1, characterized in that the polymer is at least one selected from polyracemic lactic acid, poly-L-lactic acid, poly-D-lactic acid, polyhydroxyethyl ester, polyurethane, polyamino acid, poly(diglycolic acid lactate) acid, poly-D,L-lactide, polypropylene glycol, polyglycolic acid, polyglycolic acid polylactic acid, polysalicylic anhydride ester, polytrimethylene carbonate, polycaprolactone, polyhydroxyalkanoate, polyacrylate, polysilicate, poly(β-hydroxybutyric acid) and polyethylene adipate, poly-1,3-bis(p-carboxyphenoxy)propane-sebacic acid, poerciate dimer-sebacic acid, and polyfumarate-sebacic acid.

13. The coating for a medical device according to claim 1, characterized in that the drug supported on the drug-carrying layer is at least one of an angiogenic inhibitor, an anti-inflammatory drug, an antithrombotic drug, and an anti-allergic drug.

14. The coating for a medical device according to claim 1, characterized in that the drug supported on the drug-carrying layer is at least one of paclitaxel, rapamycin, cilostazol, heparin, dexamethasone, tacrolimus, everolimus, calcium gluconate, chlorpheniramine, or cortisone.

15. The coating for a medical device according to claim 1, characterized in that the medical device is a biodegradable medical device, and the medical device is a biodegradable metallic medical device or a biodegradable nonmetallic medical device.

16. The coating for the medical device according to claim 1, characterized in that the medical device includes any of the following: a vascular stent, a heart valve, a non-vascular lumen stent, an occluder, an orthopedic implant, a dental implant, a respiratory implant, a gynecological implant, a male medical implant, a suture, or a bolt.

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