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Elastomeric copolymer coatings for implantable medical devices

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Elastomeric copolymer coatings for implantable medical devices


Implantable medical devices with elastomeric copolymer coatings are disclosed.

Browse recent Abbott Cardiovascular Systems Inc. patents - Santa Clara, CA, US
Inventor: Yunbing Wang
USPTO Applicaton #: #20120330404 - Class: 623 138 (USPTO) - 12/27/12 - Class 623 
Prosthesis (i.e., Artificial Body Members), Parts Thereof, Or Aids And Accessories Therefor > Arterial Prosthesis (i.e., Blood Vessel) >Absorbable In Natural Tissue



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The Patent Description & Claims data below is from USPTO Patent Application 20120330404, Elastomeric copolymer coatings for implantable medical devices.

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CROSS-REFERENCE TO RELATED APPLICATION

This application is a continuation of U.S. application Ser. No. 11/810,652, filed on Jun. 5, 2007, and published as U.S. Patent Application Publication No. 2008-0306592 A1, on Dec. 11, 2008, which is incorporated by reference in its entirety, including any drawings, herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to elastomeric coatings for implantable medical devices.

2. Description of the State of the Art

This invention relates to radially expandable endoprostheses, which are adapted to be implanted in a bodily lumen. An “endoprosthesis” corresponds to an artificial device that is placed inside the body. A “lumen” refers to a cavity of a tubular organ such as a blood vessel.

A stent is an example of such an endoprosthesis. Stents are generally cylindrically shaped devices, which function to hold open and sometimes expand a segment of a blood vessel or other anatomical lumen such as urinary tracts and bile ducts. Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments, stents reinforce body vessels and prevent restenosis following angioplasty in the vascular system. “Restenosis” refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through a bodily lumen to a region, such as a lesion, in a vessel that requires treatment. “Deployment” corresponds to the expanding of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into a bodily lumen, advancing the catheter in the bodily lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about a balloon disposed on the catheter. Mounting the stent typically involves compressing or crimping the stent onto the balloon. The stent is then expanded by inflating the balloon. The balloon may then be deflated and the catheter withdrawn. In the case of a self-expanding stent, the stent may be secured to the catheter via a constraining member such as a retractable sheath or a sock. When the stent is in a desired bodily location, the sheath may be withdrawn which allows the stent to self-expand.

The stent must be able to satisfy a number of mechanical requirements. First, the stent must be capable of withstanding the structural loads, namely radial compressive forces, imposed on the stent as it supports the walls of a vessel. Therefore, a stent must possess adequate radial strength. Radial strength, which is the ability of a stent to resist radial compressive forces, is due to strength and rigidity around a circumferential direction of the stent. Radial strength and rigidity, therefore, may also be described as, hoop or circumferential strength and rigidity.

Once expanded, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it, including the cyclic loading induced by the beating heart. For example, a radially directed force may tend to cause a stent to recoil inward. Generally, it is desirable to minimize recoil. In addition, the stent must possess sufficient flexibility to allow for crimping, expansion, and cyclic loading. Longitudinal flexibility is important to allow the stent to be maneuvered through a tortuous vascular path and to enable it to conform to a deployment site that may not be linear or may be subject to flexure. Finally, the stent must be biocompatible so as not to trigger any adverse vascular responses.

The structure of a stent is typically composed of scaffolding that includes a pattern or network of interconnecting structural elements often referred to in the art as struts or bar arms. The scaffolding can be formed from wires, tubes, or sheets of material rolled into a cylindrical shape. The scaffolding is designed so that the stent can be radially compressed (to allow crimping) and radially expanded (to allow deployment). A conventional stent is allowed to expand and contract through movement of individual structural elements of a pattern with respect to each other.

Furthermore, it may be desirable for a stent to be biodegradable. In many treatment applications, the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Therefore, stents fabricated from biodegradable, bioabsorbable, and/or bioerodable materials such as bioabsorbable polymers should be configured to completely erode only after the clinical need for them has ended.

Additionally, a medicated stent may be fabricated by coating the surface of either a metallic or polymeric scaffolding with a polymeric carrier that includes an active or bioactive agent or drug. Polymeric scaffolding may also serve as a carrier of an active agent or drug. Potential problems with therapeutic coatings for polymeric implantable medical devices, such as stents, include insufficient toughness, slow degradation rate, and poor adhesion.

SUMMARY

OF THE INVENTION

Certain embodiments of the present invention include an implantable medical device comprising a coating above a polymer surface of the device, the coating comprising: a block copolymer including an elastic block and an anchor block, the elastic block being a homopolymer and elastomeric at physiological conditions, the anchor block being miscible with the surface polymer.

Further embodiments of the present invention include an implantable medical device comprising a coating above a polymer surface of the device, the coating comprising: a elastomeric copolymer including elastic units and anchor units, the elastic units providing elastomeric properties to the copolymer at physiological conditions, wherein the anchor units enhance adhesion of the coating with the surface polymer, wherein the copolymer is a star block copolymer having at least three arms, the arms comprising the elastic units and the anchor units.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a view of a stent.

FIG. 2A depicts a cross-section of a stent surface with a block copolymer coating layer over a substrate.

FIG. 2B depicts a cross-section of a stent surface with a block copolymer coating layer over a polymeric layer disposed over a substrate of the stent.

FIG. 3 depicts a cross-section of a stent surface with the block-copolymer coating layer over a substrate of the stent showing an interfacial region.

FIG. 4 depicts a cross-section of a stent showing a coating material layer over a swollen surface polymer layer.

FIG. 5 depicts a polymer surface pretreated with a solvent.

FIG. 6 depicts the cross-section of a stent surface with a drug-polymer layer over a block copolymer primer layer disposed over a substrate of the stent.

DETAILED DESCRIPTION

OF THE INVENTION

Various embodiments of the present invention include an implantable medical device with a coating having an elastomeric polymer above a polymeric surface of the device. The polymeric surface may be a surface of a polymer coating disposed above a substrate that can be composed of metal, polymer, ceramic, or other suitable material. Alternatively, the polymeric surface may be a surface of a polymeric substrate or body. “Above” a surface is defined as higher than or over a surface measured along an axis normal to the surface, but not necessarily in contact with the surface.

The present invention may be applied to implantable medical devices including, but not limited to, self-expandable stents, balloon-expandable stents, stent-grafts, and grafts (e.g., aortic grafts), and generally expandable tubular devices for various bodily lumen or orifices. A stent can have a scaffolding or a substrate that includes a pattern of a plurality of interconnecting structural elements or struts. FIG. 1 depicts a view of an exemplary stent 100. Stent 100 includes a pattern with a number of interconnecting structural elements or struts 110. In general, a stent pattern is designed so that the stent can be radially compressed (crimped) and radially expanded (to allow deployment). The stresses involved during compression and expansion are generally distributed throughout various structural elements of the stent pattern. The variations in stent patterns are virtually unlimited.

In some embodiments, a stent may be fabricated by laser cutting a pattern on a tube or a sheet rolled into a tube. Representative examples of lasers that may be used include, but are not limited to, excimer, carbon dioxide, and YAG. In other embodiments, chemical etching may be used to form a pattern on a tube.

An implantable medical device can be made partially or completely from a biodegradable, bioabsorbable, biostable polymer, or a combination thereof. A polymer for use in fabricating an implantable medical device can be biostable, bioabsorbable, biodegradable or bioerodable. Biostable refers to polymers that are not biodegradable. The terms biodegradable, bioabsorbable, and bioerodable are used interchangeably and refer to polymers that are capable of being completely degraded and/or eroded when exposed to bodily fluids such as blood and can be gradually resorbed, absorbed, and/or eliminated by the body. The processes of breaking down and absorption of the polymer can be caused by, for example, hydrolysis and metabolic processes.

As indicated above, a medicated implantable medical device, such as a stent, may be fabricated by coating the surface of the device with a drug. For example, a stent can have a coating including a drug dispersed in a polymeric carrier disposed over a substrate of the stent. Such a coating layer may be formed by applying a coating material to a substrate of an implantable medical device, such as a stent. The coating material can be a polymer solution and a drug dispersed in the solution. The coating material may be applied to the stent by immersing the stent in the coating material, by spraying the material onto the stent, or by other methods known in the art. The solvent in the solution is then removed, for example, by evaporation, leaving on the stent surfaces a polymer coating impregnated with the drug.

Stents are typically subjected to stress during use. “Use” includes manufacturing, assembling (e.g., crimping a stent on balloon), delivery of a stent through a bodily lumen to a treatment site, deployment of a stent at a treatment site, and treatment after deployment. Both the underlying scaffolding or substrate and the coating experience stress that result in strain in the substrate and coating. In particular, localized portions of the stent's structure undergo substantial deformation. For example, the apex regions of bending elements 130, 140, and 150 in FIG. 1 experience relatively high stress and strain during crimping, expansion, and after expansion of the stent.

Furthermore, polymer substrates and polymer-based coatings may be particularly vulnerable to mechanical instability during use of a stent. Such mechanical instability for coatings can include fracture and detachment from a substrate, for exampling, peeling. Some polymers may be susceptible to such mechanical instability due to insufficient toughness at high deformations. Additionally, detachment of coatings may be due to poor adhesion of the polymer-based coating to the substrate or another polymer layer. Therefore, polymer-based coatings are highly susceptible to tearing or fracture, and/or detachment, especially at regions subjected to relatively high stress and strain. Thus, it is important for a polymer-based coating to (1) be tough and have a high resistance to cracking and (2) have good adhesion with an underlying layer or substrate and to have a high resistance to detachment in the range of deformations that occur during crimping, during deployment of a stent, and after deployment.

As indicated above, a device may be composed in whole or in part of materials that degrade, erode, or disintegrate through exposure to physiological conditions within the body until the treatment regimen is completed. The device may be configured to disintegrate and disappear from the region of implantation once treatment is completed. The device may disintegrate by one or more mechanisms including, but not limited to, dissolution and chemical breakdown. The duration of a treatment period depends on the bodily disorder that is being treated. For illustrative purposes only, in treatment of coronary heart disease involving use of stents in diseased vessels, the duration can be in a range from about a month to a few years. However, the duration is typically in a range from about six to twelve months. Thus, it is desirable for polymer-based coatings and substrates of an implantable medical device, such as a stent, to have a degradation time at or near the duration of treatment. Degradation time refers to the time for an implantable medical device to substantially or completely erode away from an implant site.

Embodiments of the present invention can include an elastomeric polymer coating disposed over a polymer surface of a device, such as a stent scaffolding. In certain embodiments, the coating can be disposed directly over the surface of a polymer substrate of a device. FIG. 2A depicts a cross-section of a stent surface with an elastomeric polymer coating layer 210 over a substrate 200. In the embodiment shown in FIG. 2A, elastomeric polymer coating layer 210 includes a drug 220 dispersed in an elastomeric polymer 230. The substrate can be composed of a bioabsorbable polymer.

In other embodiments of the present invention, the elastomeric polymer coating can be over a polymer coating layer that is disposed over a substrate. FIG. 2B depicts a cross-section of a substrate 240 of a stent with a polymeric layer 250 disposed over substrate 240. An elastomeric polymer coating layer 260 is disposed over polymeric layer 250. Coating layer 260 includes a drug 270 dispersed within an elastomeric polymer 280. Polymeric layer 250 can be a primer layer for improving the adhesion of drug-polymer layer 260 to substrate 240. In the embodiment of FIG. 2B, substrate 240 can be metallic, polymeric, ceramic, or other suitable material.

In certain embodiments of the present invention, the elastomeric polymer coating can include a block copolymer having an elastic block and an anchor block. In such embodiments, the elastic block is a homopolymer that exhibits elastomeric or rubbery behavior at physiological conditions. In addition, the anchor block is miscible with the surface polymer and enhances the adhesion of the block copolymer coating with the surface polymer. In some embodiments, the elastic block, the anchor block, or both can be bioabsorbable polymers. In certain embodiments, all or a majority of the coating may be the block copolymer. Additionally, the coating can be a therapeutic layer with an active agent or drug mixed or dispersed within the block copolymer.

As mentioned above, the block copolymer coating exhibits rubbery or elastomeric behavior at physiological conditions. An “elastomeric” or “rubbery” polymer refers to a polymer that exhibits elastic deformation through all or most of a range of deformation. Physiological conditions include, but are not limited to, human body temperature, approximately 37° C. The elastic block of the block copolymer is an elastomeric or rubbery polymer that allows or provides the elastomeric or rubbery properties of the coating. Such elastomeric properties provide the coating with a high fracture toughness during use of a device such as a stent.

In some embodiments, the elastic blocks can have a glass transition temperature (Tg) below body temperature. Additionally, the block copolymer may be completely or substantially amorphous. Exemplary biodegradable polymers that are elastomeric or rubbery at physiological conditions include, but are not limited to, polycaprolactone (PCL), poly(tetramethyl carbonate) (PTMC), poly(4-hydroxy butyrate) (PHB), and polydioxanone (PDO).

As discussed above, the anchor block of the block copolymer can be miscible with the surface polymer. In one embodiment, the anchor block can have the same chemical composition as the surface polymer. Alternatively, the anchor block can have a chemical composition different from the surface polymer, but similar enough so that the anchor block is miscible with the surface polymer. In an exemplary embodiment, the block copolymer can have a PLLA anchor block and be disposed over a PLLA surface, which can be the surface of a PLLA substrate. In another exemplary embodiment, the block copolymer can have a PLLA anchor block and be disposed over a poly(L-lactide-co-glycolide) (LPLG) surface, which can be the surface of an LPLG substrate.

In certain embodiments, the anchor block can be a random copolymer. In such embodiments, the composition of the anchor block copolymer of the block copolymer coating can be selected so that the anchor block is miscible with the surface polymer. In addition, the units of the copolymer can be selected to adjust the degradation rate of the block copolymer. In one embodiment, the anchor block can include units that are more hydrolytically active or hydrophilic than other units to increase the degradation rate of the coating. In an exemplary embodiment, the anchor block can be LPLG. In such an embodiment, the surface polymer can be an LPLG copolymer. The composition of LLA and GA in the anchor block can be adjusted so that the LPLG anchor block is miscible with the LPLG surface polymer. In some embodiments, the surface polymer can be a copolymer having a high percentage of LLA units, for example, at least 60 wt %, 70 wt %, or 80 wt % LLA units.

In additional embodiments, the block copolymer can additionally include a fast degrading block that is selected to increases the degradation rate of the block copolymer coating. In some embodiment, the fast degrading blocks can be glassy at physiological conditions or have a Tg above body temperature. Additionally or alternatively, the fast degrading blocks can be immiscible with the surface polymer.

In some embodiments, the fast degrading block may have a higher degradation rate than the anchor block, the elastic block, or both. The fast degrading block may be composed of units that are more hydrophilic or more hydrolytically active than the elastic block or the anchor block. Additionally, fast degrading block may have acidic and hydrophilic degradation products. Since the rate of the hydrolysis reaction tends to increase as the pH decreases, acidic degradation products can increase the degradation rate of the block copolymer coating. Glycolide (GA) units, for example, have acidic degradation products which can increase the degradation rate of the coating. Exemplary fast degrading blocks can include poly(glycolide) (PGA) and LPLG that may not be miscible with a surface polymer.

In some embodiments, the toughness of the block copolymer coating can be adjusted by increasing or decreasing the weight percent of elastic blocks. As the weight percent of elastic blocks increases, the block copolymer can become more flexible and tougher. For example, for a PCL-b-PLLA coating, as the weight percent of PCL increases, the block copolymer becomes more flexible and tougher. The composition of the elastic blocks of the block copolymer can be greater than 20 wt %, 30 wt %, 40 wt %, 50 wt %, 60 wt %, 70 wt %, 80 wt %, 90 wt %, or greater than 90 wt % of the block copolymer.

In exemplary embodiments, the molecular weight of the elastic blocks can be between 20 kg/mol and 150 kg/mol, or greater than 150 kg/mol. The molecular weight of the anchor blocks can be between 20 kg/mol and 150 kg/mol, or greater than 150 kg/mol. The relative weight percent of the elastic blocks and the anchor blocks can be between 1:5 and 10:1.

Additionally, in other embodiments, the degradation rate of the coating can be adjusted by increasing or decreasing the weight percent of fast degrading blocks. The degradation rate of the coating can be increased by increasing the weight percent of fast degrading blocks. For example, the weight percent of PGA in a PLLA-b-PGA-b-PDO block copolymer can be increased to increase the degradation rate of the polymer.

Embodiments of the block copolymer of the elastomeric coating can have two or more blocks. The block copolymer can be a diblock, triblock, tetrablock, pentablock, etc. copolymer. Diblock copolymers can include, for example, PLLA-b-PDO, PLLA-b-PCL, and PLLA-b-PTMC. Exemplary triblock copolymers include PLLA-b-PDO-b-PLLA, PLLA-b-PCL-PLLA, and PLLA-b-PTMC-b-PLLA, PLLA-b-PGA-b-PDO, etc. Such block copolymers may be suitable as coatings over a PLLA or LPLG surface.

In some embodiments, the block copolymer can be a branched polymer which corresponds to a polymer with “side chains.” Branched polymers include, for example, hyperbranched-like polymers, comb-like polymers, star polymers, dendrimer-like star polymers, and dendrimers. A star polymer refers to a polymer having at least three chains or arms radiating outward from a common center. A dendritic polymer is a branched polymer resembling a tree-like structure. A comb structure corresponds to a linear polymer segment or backbone having a plurality of side chains extending outward from a position along the linear segment. In such embodiments, a block copolymer can be a branched polymer with at least one branch that is an elastic block and at least one branch that is an anchor block. The branched block copolymer can further include at least one branch that is a fast degrading block.

In these embodiments, the block copolymer can be a star block copolymer having at least three arms or branches with at least one arm being an elastic block and at least one arm being an anchor block. The star block copolymer can further include at least one arm that is a fast degrading block.

In further embodiments of the present invention, the elastomeric coating above a surface polymer of an implantable medical device can include a random copolymer with elastic units and anchor units. In such embodiments, the elastic units provide elastomeric or rubbery properties at physiological conditions to the random copolymer. “Elastic units,” refer to monomer units that form elastic or rubbery polymers at physiological conditions. Exemplary elastic units can include, but are not limited to, caprolactone (CL), tetramethyl carbonate (TMC), 4-hydroxy butyrate (HB), and dioxanone (DO). In addition, the anchor units enhance the adhesion of the random copolymer coating with the surface polymer. In some embodiments, the elastic units, the anchor units, or both can be bioabsorbable. In certain embodiments, all or a majority of the coating may be the random copolymer. Additionally, the coating can be a therapeutic layer with an active agent or drug mixed or dispersed within the random copolymer.

Alternatively, the elastomeric copolymer can be an alternating copolymer with elastic units and anchor units alternating along the polymer chain. In addition, the elastomeric copolymer can include more than one type of elastic unit and more than one type of anchor unit.

The anchor units of the random copolymer are the same as at least one unit in the surface polymer. Additionally, the anchor units can be miscible with the surface polymer. The anchor units can allow portions of segments of the random copolymer to be miscible with the surface polymer. The degree of adhesion can be increased by increasing the weight percent of the anchor units in the copolymer. In an exemplary embodiment, the surface polymer can be a crystalline or semicrystalline polymer. In such embodiments, the anchor units can be units of such a crystalline or semicrystalline polymer. In an exemplary embodiment, the random copolymer can have an LLA anchor units and be disposed over a PLLA surface, which can be the surface of a PLLA substrate. In another exemplary embodiment, the random copolymer can have a LLA, GA, or both LLA and GA anchor units and be disposed over an LPLG surface, which can be the surface of a LPLG substrate.

In additional embodiments, the random copolymer can additionally include fast degrading units, alternatively or additionally to anchor units, that are selected to increase the degradation rate of the random copolymer coating. The fast degrading units can be more hydrophilic or more hydrolytically active than the elastic units or the anchor units. Additionally, fast degrading blocks may have acidic and hydrophilic degradation products. The fast degrading units can be glassy at physiological conditions or can be different from units of the surface polymer. In an exemplary embodiment, GA units are fast degrading units in a random copolymer coating disposed over a PLLA surface polymer.

In some embodiments, the elastomeric copolymer can be a random or alternating copolymer of elastic units and fast degrading units. Alternatively, the copolymer can be a random or alternating copolymer of elastic units, anchor units, and fast degrading units.

Exemplary random copolymer coatings include PLLA-co-PDO, PLLA-co-PCL, PLLA-co-PTMC, PLLA-co-PDO-co-PTMC, PLLA-co-PGA-co-PDO, PLLA-co-PGA-co-PCL, PLLA-co-PGA-co-PTMC, etc. Such block copolymers may be suitable as coatings over a PLLA or LPLG surface.

In some embodiments, the random copolymer can be a branched polymer, including, for example, hyperbranched-like polymers, comb-like polymers, star polymers, dendrimer-like star polymers, and dendrimers. In such embodiments, the random copolymer can be a random branched copolymer having branches including elastic units and anchor units. The branches of branched polymers can further include fast degrading units. In exemplary embodiments, the random copolymer can be a random star copolymer having at least three arms or branches with at least one arm including elastic units and anchor units. The arms can also include fast degrading units.

Embodiments of the elastomeric polymer coating of the present invention can be applied to a polymer surface so that at least some of the elastomeric polymer is mixed with the surface polymer. In particular, at least the anchor blocks of the block copolymer coating can be mixed within the surface polymer. Alternatively, segments of the random copolymer of the random copolymer coating that include anchor units can be mixed with the surface polymer. It is believed that an interfacial region between the coating and the surface polymer can form with elastomeric polymer mixed with surface polymer. The anchor blocks of the block copolymer or the anchor units of the random copolymer can act as a compatibilizer that strengthens the bond between the coating and the coated surface. The interfacial region can enhance the adhesion of the elastomeric polymer coating to the polymer substrate or a polymer surface layer, in general.

FIG. 3 depicts a cross-section of a stent surface with an elastomeric copolymer coating layer 310 over a substrate 300. Coating layer 310 can be applied to form an interfacial region 340 which can include anchor blocks or random copolymer segments including anchor units mixed with substrate polymer. A drug 320 can be mixed or dispersed within coating layer 310 and interfacial region 340. A thickness Ti of interfacial region 340 can be varied depending on coating application processing parameters.

The enhanced adhesion can allow the use of a tough, high fracture resistant coating that may otherwise have poor adhesion to a polymer substrate of a device. The polymer material for a substrate of a device, such as a stent, may be selected primarily on the basis of strength and stiffness so that the stent substrate can provide support for a lumen. Such substrate polymers can be crystalline or semi-crystalline polymers that are glassy or have a Tg above body temperature. Tough, elastomeric polymers may not necessarily have good adhesion with such a substrate. Embodiments of the block copolymer or random copolymer disclosed herein allow the use of a tough, high fracture resistant coating over a glassy substrate. Such glassy substrate polymers include PLLA and LPLG.

In some embodiments, the block copolymer coating can include a dispersed polymer phase. In such embodiments, the anchor block can have a high enough molecular weight that a dispersed anchor block phase is formed within an elastomeric phase composed of the elastic blocks. In these embodiments, the anchor block can be a crystalline or semicrystalline polymer. The dispersed phase can be crystalline or semi-crystalline polymer regions that are dispersed within an amorphous elastomeric phase. The crystalline regions can be used to modify the delivery rate of a dispersed drug from the coating. The crystalline regions tend to increase the delivery rate of drug from the coating.

Embodiments of the elastomeric polymers disclosed herein can be formed by solution-based polymerization. Other methods of forming the elastomeric polymers are also possible, such as, without limitation, melt phase polymerization.

Some embodiments of the solution polymerization involve forming the elastic blocks first and then the anchor blocks. In such embodiments, a solution is prepared including the elastic units for the elastic blocks, an appropriate solvent, an appropriate initiator, and catalyst. The elastic blocks are formed in the solution from the monomers. The anchor block units for the anchor block and catalyst are then added to the solution to form anchor blocks that are bonded to the elastic blocks. The elastomeric block copolymer can be removed from the solution through precipitation in a non-solvent of the elastomeric block copolymer. The solvent(s) for the reaction mixture can be selected so that the elastic blocks formed are soluble in the solvent(s) to allow the elastic blocks to copolymerize with anchor blocks in solution.

For example, to prepare PDO-b-PLLA diblock copolymer, PDO elastic blocks are formed in a solution containing DO monomers, a dodecanol initiator, and stannous octoate catalyst in a toluene solvent. L-lactide monomers are then added to the solution. The L-lactide monomers react with PDO to form PDO-b-PLLA. The solution can then be added to methanol, which is a non-solvent for the formed block copolymer, to precipitate the elastomeric block copolymer from solution. Other embodiments of the solution polymerization involve forming the anchor blocks first and then the elastic blocks.

In other embodiments of solution polymerization, elastomeric block copolymers can be formed by reacting elastic blocks swollen with a solvent that contain anchor block monomer units. The elastic blocks are swollen by a solvent after they are formed so that they can react with anchor block monomer units. One of skill in the art can select a solvent that swells but does not dissolve the elastic blocks.

The elastomeric random copolymer can be prepared by solution polymerization by preparing a solution including the elastic units, anchor units, optionally fast degrading units, an appropriate solvent, an appropriate initiator, and catalyst. The mixture is allowed to react to form the elastomeric random copolymer. The elastomeric random copolymer can be removed from the solution through precipitation in a non-solvent of the elastomeric random copolymer.

For example, to prepare a PLLA-co-PCL-co-PDO random copolymer, a solution is formed containing DO units, CL units, and LLA units, a dodecanol initiator, and stannous octoate catalyst in a toluene solvent. The monomers react to form the random copolymer. The solution can then be added to methanol, which is a non-solvent for the formed random copolymer, to precipitate the elastomeric random copolymer from solution.

In one embodiment, the solvent for use in synthesizing the elastomeric block copolymer is devoid of alcohol functional groups. Such alcohol groups may act as initiators for chain growth in the polymer. Solvents used to synthesize the elastomeric block copolymer include, but are not limited to, chloroform, toluene, xylene, and cyclohexane.

Elastomeric star block copolymers and random copolymers can be synthesized according to the schemes described above by using an appropriate initiator. In one embodiment, pentaerythritol can be used as an initiator to synthesize star polymers.

Embodiments of the elastomeric polymer coating of the present invention may be formed over an implantable medical device, such as a stent, by applying a coating material to a polymer surface of the device. The coating material can be a solution including the elastomeric copolymer. The solution can further include an active agent or drug dissolved in a solvent. As discussed above, the coating material may be applied to the stent by immersing the device in the coating material, by spraying the composition onto the device, or by other methods known in the art. The solvent in the applied solution is removed, leaving on the device surfaces the elastomeric polymer coating and optionally drug dispersed within the polymer.

Drying or solvent removal can be performed by allowing the solvent to evaporate at room or ambient temperature. Depending on the volatility of the particular solvent employed, the solvent can evaporate essentially upon contact with the stent. Alternatively, the solvent can be removed by subjecting the coated stent to various drying processes. Drying time can be decreased to increase manufacturing throughput by heating the coated stent. For example, removal of the solvent can be induced by baking the stent in an oven at a mild temperature (e.g., 50° C.) for a suitable duration of time (e.g., 2-4 hours) or by the application of warm air. In an embodiment, a substantial portion of solvent removed may correspond to less than 5%, 1%, or more narrowly, less than 0.5% of solvent remaining after drying. Depositing a coating of a desired thickness in a single coating stage can result in an undesirably nonuniform surface structure and/or coating defects. Therefore, a coating process can involve multiple repetitions of application, for example, by spraying a plurality of layers.

In some embodiments, the solvent of the coating material is also a solvent for the surface polymer on which the coating material is applied. Specifically, a “solvent” for a given polymer can be defined as a substance capable of dissolving or dispersing the polymer or capable of at least partially dissolving or dispersing the polymer to form a uniformly dispersed mixture at the molecular- or ionic-size level. The solvent should be capable of dissolving at least 0.1 mg of the polymer in 1 ml of the solvent, and more narrowly 0.5 mg in 1 ml at ambient temperature and ambient pressure. The solvent in the coating material can dissolve at least a portion of the surface polymer upon application of the coating material to the polymer surface.

Due to dissolution of a portion of the surface polymer, the coating material near the surface of the surface polymer includes dissolved surface polymer in addition to the elastomeric polymer from the coating material. It is believed that upon removal of the solvent, an interfacial region, as depicted in FIG. 3, is formed that includes anchor blocks of the block copolymer or segments of the random copolymer mixed with surface polymer. This interfacial region can be formed due to the miscibility of the surface polymer with the anchor blocks or segments including anchor units.

In other embodiments, the solvent in the coating material can be capable of swelling the surface polymer, but is incapable or substantially incapable of dissolving the surface polymer. A solvent that is capable of swelling the surface polymer and is incapable or substantially incapable of dissolving the polymer is understood to mean a sample of the surface polymer swells when immersed in the solvent and the swollen sample of the surface polymer remains in the solvent with a negligible loss of mass for an indefinite period of time at conditions of ambient temperature and pressure.

Solvents for polymers can be found in standard texts (e.g., see Fuchs, in Polymer Handbook, 3rd Edition and Deasy, Microencapsulation and Related Drug Processes, 1984, Marcel Dekker, Inc., New York.) The ability of a polymer to swell and to dissolve in a solvent can be estimated using the Cohesive Energy Density Concept (CED) and related solubility parameter values as discussed by Deasy and can be found in detail in the article by Grulke in Polymer Handbook.

FIG. 4 depicts a cross-section of a stent showing a coating material layer 400 over a swollen surface polymer layer 410. Swollen surface polymer layer 410 is over unswollen polymer coating layer or polymer substrate 420. As indicated above, unswollen surface polymer 420 can either be a substrate of the stent or a polymeric coating over a stent substrate. As shown, swollen surface polymer layer 410 has a thickness Ts. Due to swelling of the surface polymer in swollen surface polymer layer 410, it is believed that anchor blocks or segments containing anchor units of the elastomeric polymer in coating material layer 400 penetrate into or mix with the surface polymer in swollen polymer layer 410 prior to removal of the solvent. Upon removal of the solvent, a coating layer is formed over substrate 420.

In some embodiments, a polymeric substrate or polymeric surface coating layer can be pretreated with a solvent that dissolves or swells the surface polymer prior to applying a coating material. FIG. 5 depicts a layer 510 over a substrate or coating layer 500. Layer 510 can be a dissolved layer of surface polymer or a swollen layer of surface polymer. Following pretreatment, the coating material can be applied over the pretreated surface.

In certain embodiments, the coating material solvent is different from the pretreatment solvent. The use of a different solvent for the coating material and the pretreating can provide a degree of flexibility to the coating process. Generally, a treatment with a medicated stent may require a particular drug coating on a coating of a medicated stent. A drug may have an undesirably low or negligible solubility in a selected group of solvents that can dissolve or swell the surface polymer. Thus, a drug coating formed using such a solvent can have an undesirably low concentration of drug. A suitable pretreatment solvent can be used to dissolve or swell the surface polymer and a different solvent can be used as a coating solvent, in which the drug has an acceptable solubility. In general, a required solubility of a drug in a coating solvent is determined by the drug loading required of a particular treatment regimen. Specifically, it is desirable for a drug to have solubility of at least 1 wt % in a solvent for use as a coating material solvent for forming a drug-polymer layer on a stent.

In other embodiments, an elastomeric polymer coating can be a primer layer over a polymer substrate or coating layer. The elastomeric polymer coating can act as a primer layer for a drug-polymer coating layer over the primer layer. The elastomeric polymer primer layer may be formed above a polymeric surface, as described above. The primer coating material can include an elastomeric polymer dissolved in a solvent that can dissolve or swell the surface polymer. A drug-polymer layer can then be formed over the elastomeric polymer primer layer. The drug coating material may include a polymer that is different from the elastomeric polymer and a solvent that is different from the primer coating material solvent. FIG. 6 depicts a drug layer 650 over primer coating layer 630. Drug layer 650 includes a drug 660 mixed or dispersed within a polymer 670. An interfacial layer 640, discussed above, includes anchor blocks or segments including anchor units and surface polymer.

In general, representative examples of polymers that may be used to fabricated a substrate of and coatings for an implantable device include, but are not limited to, poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxyvalerate), poly(lactide-co-glycolide), poly(hydroxybutyrate), poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride, poly(glycolic acid), poly(glycolide), poly(L-lactic acid), poly(L-lactide), poly(D,L-lactic acid), poly(L-lactide-co-glycolide); poly(D,L-lactide), poly(caprolactone), poly(trimethylene carbonate), polyethylene amide, polyethylene acrylate, poly(glycolic acid-co-trimethylene carbonate), co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes, biomolecules (such as fibrin, fibrinogen, cellulose, starch, collagen and hyaluronic acid), polyurethanes, silicones, polyesters, polyolefins, polyisobutylene and ethylene-alphaolefin copolymers, acrylic polymers and copolymers other than polyacrylates, vinyl halide polymers and copolymers (such as polyvinyl chloride), polyvinyl ethers (such as polyvinyl methyl ether), polyvinylidene halides (such as polyvinylidene chloride), polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics (such as polystyrene), polyvinyl esters (such as polyvinyl acetate), acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon 66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides, polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, cellulose acetate, cellulose butyrate, cellulose acetate butyrate, cellophane, cellulose nitrate, cellulose propionate, cellulose ethers, and carboxymethyl cellulose.

Additional representative examples of polymers that may be especially well suited for use in embodiments of the present invention include ethylene vinyl alcohol copolymer (commonly known by the generic name EVOH or by the trade name EVAL), poly(butyl methacrylate), poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508, available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidene fluoride (otherwise known as KYNAR, available from ATOFINA Chemicals, Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethylene glycol.

For the purposes of the present invention, the following terms and definitions apply:

For the purposes of the present invention, the following terms and definitions apply:

The “glass transition temperature,” Tg, is the temperature at which the amorphous domains of a polymer change from a brittle vitreous state to a solid deformable, ductile, or rubbery state at atmospheric pressure. In other words, the Tg corresponds to the temperature where the onset of segmental motion in the chains of the polymer occurs. When an amorphous or semicrystalline polymer is exposed to an increasing temperature, the coefficient of expansion and the heat capacity of the polymer both increase as the temperature is raised, indicating increased molecular motion. As the temperature is raised the actual molecular volume in the sample remains constant, and so a higher coefficient of expansion points to an increase in free volume associated with the system and therefore increased freedom for the molecules to move. The increasing heat capacity corresponds to an increase in heat dissipation through movement. Tg of a given polymer can be dependent on the heating rate and can be influenced by the thermal history of the polymer. Furthermore, the chemical structure of the polymer heavily influences the glass transition.

“Stress” refers to force per unit area, as in the force acting through a small area within a plane. Stress can be divided into components, normal and parallel to the plane, called normal stress and shear stress, respectively. True stress denotes the stress where force and area are measured at the same time. Conventional stress, as applied to tension and compression tests, is force divided by the original gauge length.

“Strength” refers to the maximum stress along an axis which a material will withstand prior to fracture. The ultimate strength is calculated from the maximum load applied during the test divided by the original cross-sectional area.

“Strain” refers to the amount of elongation or compression that occurs in a material at a given stress or load.

“Elongation” may be defined as the increase in length in a material which occurs when subjected to stress. It is typically expressed as a percentage of the original length.

“Toughness” is the amount of energy absorbed prior to fracture, or equivalently, the amount of work required to fracture a material. One measure of toughness is the area under a stress-strain curve from zero strain to the strain at fracture. Thus, a brittle material tends to have a relatively low toughness.

Drugs or therapeutic active agent(s) can include anti-inflammatories, antiproliferatives, and other bioactive agents.



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stats Patent Info
Application #
US 20120330404 A1
Publish Date
12/27/2012
Document #
13598465
File Date
08/29/2012
USPTO Class
623/138
Other USPTO Classes
International Class
61F2/82
Drawings
3


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Prosthesis (i.e., Artificial Body Members), Parts Thereof, Or Aids And Accessories Therefor   Arterial Prosthesis (i.e., Blood Vessel)   Absorbable In Natural Tissue