WO2015053934A1

WO2015053934A1 - Implantable medical devices including a polyether-polyurethane substrate with improved biostability, and methods

Info

Publication number: WO2015053934A1
Application number: PCT/US2014/056957
Authority: WO
Inventors: Kimberly A. Chaffin; Xiangji Chen; Matthew Jolley
Original assignee: Medtronic, Inc.
Priority date: 2013-10-10
Filing date: 2014-09-23
Publication date: 2015-04-16

Abstract

An implantable medical device including a polyether-polyurethane substrate and a more oxidatively-stable polyurethane forming a protective layer directly disposed on at least one surface of the polyether-polyurethane substrate.

Description

Implantable Medical Devices Including a Polyether-Polyurethane Substrate with Improved Biostability, and

Methods Cross-reference to Related Applications

The present application claims priority to U.S. Provisional Application Serial No. 61 /889,170, filed October 10, 2013, which is incorporated herein by reference in its entirety. Background

There is a need for soft flexible polymers and constructs thereof with improved biostability for implantable medical devices, such as electrical lead insulation, drug infusion catheters, artificial polymer discs, etc. Widely adopted soft 80A durometer polyether-polyurethanes (e.g., those available under the tradenames PELLETHANE, ELASTHANE, TECOTHANE) have desirable hydrolytic stability, mechanical properties and processing characteristics which have led to their adoption in the implantable medical device industry. However, these materials are susceptible to oxidation, which is generally thought to occur from oxygen-based free radicals, originating from macrophages or foreign body giant cells or catalyzed by metal ions. These radicals can extract a hydrogen atom from the methylene group adjacent to the ether functionality in the soft segment, resulting in either chain scission or crosslinking. The oxidative processes observed in vivo with polyurethanes are commonly referred to as Environment Stress Cracking (ESC) and Metal Ion

Oxidation (MIO).

Over-coating metal lead conductors with fluoropolymers has successfully been employed to mitigate MIO. However, with ESC, cell-mediated oxygen-based free radicals are very reactive and consequently react very quickly at the surface prior to diffusion and reaction in the bulk. This results in a heterogeneous oxidative attack that takes place primarily at the surface of the polyether-polyurethane giving rise to the formation of surface crazing and micro-cracks. Over time, these cracks can readily propagate through the bulk under typical in vivo loading conditions (flexing, bending, stretching, fatigue, abrasion, etc.), leading to bulk mechanical failure, i.e., breaches. Furthermore, surface cracks can expose the underlying material to the oxidizing environment leading to an apparent acceleration of this process.

Since the observations back in the early 1980’s of oxidative degradation of soft polyether-polyurethanes, the implantable medical device industry has continuously explored soft 80A polyurethanes with improved oxidative stability, such as

poly(dimethyl siloxane)-polyurethanes (e.g., those available under the trade names ELAST-EON and PURSIL) and polycarbonate-polyurethanes (e.g., that available under the tradename BIONATE). However, despite their well-documented improved oxidative stability, for various reasons these materials have not been widely adopted. Polycarbonate-polyurethanes showed superior oxidation stability compared to polyether-polyurethane; however, they are known to be susceptible to hydrolysis of the carbonate group in the soft segment. PDMS-polyurethanes have recently been proven to be susceptible to hydrolysis resulting in decreasing mechanical strength (Chaffin et al., Macromolecules, 45 (22), 9110-9120 (2012)). It should be noted that hydrolysis, unlike oxidation, typically occurs throughout the polymer as water can readily diffuse into the bulk and consequently affects bulk mechanical properties, whereas cell-mediated oxidation takes place primarily at the surface. Further, the relative timescales of hydrolysis and oxidation in vivo can be different. For example, the cell-mediated inflammatory response that results in oxidation is typically most acute within the first few weeks and months of implantation, which recruits the celluar response that is the source of the oxidative free radicals. After this initial period, the implanted device typically becomes encapsulated by host tissue growth (fibrotic capsule), reducing the inflammation. Whereas hydrolysis would be expected to occur more uniformly throughout the implant life, because water is pervasive and quickly reaches equilibrium concentrations within the polymer, hydrolysis remains relatively constant over time.

Recently another class of promising implantable polyurethanes,

polyisobutylene-polyurethanes, has been developed that has shown superior oxidative stability compared to polyether-polyurethanes (U.S. Pat. No. 8,501 ,831 )^. However, polyisobutylene-polyurethanes typically have inferior mechanical properties compared with polyether-polyurethanes. Also, their synthesis and scale-up is costly and complex. Thus, there is still a need for materials that have the desirable hydrolytic stability and mechanical properties of soft polyether-polyurethanes, while improving their oxidative resistance. Summary

The present disclosure provides constructs that retain the desirable hydrolytic stability and mechanical properties of soft polyether-polyurethanes, while improving their oxidative resistance. This is accomplished by employing a more oxidatively- stable polyurethane as a protective, typically surface, layer. The protective layer prevents the initiation and propagation of cracks from oxidation that could eventually lead to breaching.

In one embodiment, there is provided an implantable medical device that includes: a substrate including a polyether-polyurethane; and a protective layer disposed directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer includes a more oxidatively-stable polyurethane; wherein the more oxidatively-stable polyurethane has a Shore A hardness of 50 to 95. In certain embodiments, the polyether-polyurethane has a Shore A hardness of 50 to 95, and in certain embodiments, the hardness of the polyether-polyurethane is outside this range, typically higher than this range (e.g., Shore D hardness of 55, which correlates to a Shore A hardness of 99).

In another embodiment, there is provided an implantable medical device that includes: a polyether-polyurethane layer; and a protective layer disposed directly on at least one surface of the polyether-polyurethane layer, wherein the protective layer includes a PDMS-polyurethane.

In yet another embodiment, there is provided a method of improving the oxidative stability of an implantable medical device. The method includes: providing a substrate that includes a polyether-polyurethane; and applying a protective layer directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer includes a more oxidatively-stable polyurethane.

As used herein, an“implantable medical device” is a device that has surfaces that contact tissue, bone, blood or other bodily fluids in the course of their operation, which are implanted into a body, and remain in the body for an extended period (e.g., greater than 28 days). Examples include vascular grafts, stents, pacemaker leads, heart valves, hemodialysis catheters, and the like, that are implanted in blood vessels or in the heart. Further examples include electrical stimulation leads, drug infusion catheters, peritoneal catheters, and the like, that are implanted in various types of tissue or bone. An“implantable medical device” can also include devices for temporary intravascular use such as catheters, guide wires, and the like, which are placed into the blood vessels or the heart for purposes of monitoring or repair.

As used herein, the phrase“more oxidatively-stable polyurethanes” refers to polyurethane multi-block polymers that show increased stability, compared with polyether-polyurethanes (particularly the polyether-polyurethane on which the more oxidatively-stable polyurethane is disposed), to chemical and physical changes when exposed to an oxidative environment. Oxidative stability could be measured by physical changes including surface pitting or cracking, mechanical property changes, or chemical changes including molecular weight changes, changes in chemical composition, etc.

As used herein,“soft segment” is a portion of the polyurethane and has a glass transition temperature (Tg) that is below body temperature.“PDMS- polyurethane” is a polyurethane where the soft segment contains

polydimethylsiloxane (PDMS).“Polycarbonate-polyurethane” is a polyurethane where the soft segment contains polycarbonate.“Polyisobutylene-polyurethane” is a polyurethane where the soft segment contains polyisobutylene.

The terms“comprises” and variations thereof do not have a limiting meaning where these terms appear in the description and claims. Such terms will be

understood to imply the inclusion of a stated step or element or group of steps or elements but not the exclusion of any other step or element or group of steps or elements. By“consisting of” is meant including, and limited to, whatever follows the phrase“consisting of.” Thus, the phrase“consisting of” indicates that the listed elements are required or mandatory, and that no other elements may be present. By “consisting essentially of” is meant including any elements listed after the phrase, and limited to other elements that do not interfere with or contribute to the activity or action specified in the disclosure for the listed elements. Thus, the phrase

“consisting essentially of” indicates that the listed elements are required or mandatory, but that other elements are optional and may or may not be present depending upon whether or not they materially affect the activity or action of the listed elements.

The words“preferred” and“preferably” refer to embodiments of the disclosure that may afford certain benefits, under certain circumstances. However, other embodiments may also be preferred, under the same or other circumstances.

Furthermore, the recitation of one or more preferred embodiments does not imply that other embodiments are not useful, and is not intended to exclude other embodiments from the scope of the disclosure.

In this application, terms such as“a,”“an,” and“the” are not intended to refer to only a singular entity, but include the general class of which a specific example may be used for illustration. The terms“a,”“an,” and“the” are used interchangeably with the term“at least one.”

The phrases“at least one of” and“comprises at least one of” followed by a list refers to any one of the items in the list and any combination of two or more items in the list.

As used herein, the term“or” is generally employed in its usual sense including“and/or” unless the content clearly dictates otherwise. The term“and/or” means one or all of the listed elements or a combination of any two or more of the listed elements.

Also herein, all numbers are assumed to be modified by the term“about” and preferably by the term“exactly.” As used herein in connection with a measured quantity, the term“about” refers to that variation in the measured quantity as would be expected by the skilled artisan making the measurement and exercising a level of care commensurate with the objective of the measurement and the precision of the measuring equipment used.

Also herein, the recitations of numerical ranges by endpoints include all numbers subsumed within that range as well as the endpoints (e.g., 1 to 5 includes 1 , 1.5, 2, 2.75, 3, 3.80, 4, 5, etc.). Herein,“up to” a number (e.g., up to 50) includes the number (e.g., 50).

The term“room temperature” refers to a temperature of 20^oC to 25^oC or 22^oC to 25^oC. Reference throughout this specification to“one embodiment,”“an

embodiment,”“certain embodiments,” or“some embodiments,” etc., means that a particular feature, configuration, composition, or characteristic described in

connection with the embodiment is included in at least one embodiment of the disclosure. Thus, the appearances of such phrases in various places throughout this specification are not necessarily referring to the same embodiment of the disclosure. Furthermore, the particular features, configurations, compositions, or characteristics may be combined in any suitable manner in one or more embodiments.

The above summary of the present disclosure is not intended to describe each disclosed embodiment or every implementation of the present disclosure. The description that follows more particularly exemplifies illustrative embodiments. In several places throughout the application, guidance is provided through lists of examples, which examples can be used in various combinations. In each instance, the recited list serves only as a representative group and should not be interpreted as an exclusive list. Drawings

The disclosure may be more completely understood in connection with the following drawings.

Figure 1. Schematics of (Figure 1 A) a cardiac lead and (Figrue 1 B) a deep brain stimulation lead (top view).

Figure 2. Schematic of a drug infusion catheter ((Figure 2A) for brain; (Figure 2B) for spine).

Figure 3. Schematic of a cross-section of a tube of a polyether-polyurethane (left side) with a protective layer only on the outside of the tube, and (right side) protection layers on both the inside and outside of the tube.

Figure 4. Renderings of SEM images for PELLETHANE 80A (Figure 4A, Figure 4B) and ELAST-EON 2A (Figure 4C, Figure 4D) after exposure to hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M) for an 8-week period (Figure 4A and Figure 4C, 400X for both) and a 16-week period (40X for PELLETHANE 80A (Figure 4B) and 400X for ELAST-EON 2A (Figure 4D)). Figure 5. Tensile stress-strain curves after 16-week exposure of hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M), the ELAST-EON 2A (E2A)-coated ELASTHANE 80A (“Coated E80A”) showed increased retention of tensile strength (solid lines) and elongation compared with the bare ELASTHANE 80A polyether-polyurethane (“Bare E80A”) (dashed lines).

Figure 6. Mechanical properties (tensile) and molecular weight data (weight averaged molecular weight) of ELASTHANE 80A polyether-polyurethane (“E80A”) and polyisobutylene-polyurethane/ELASTHANE 80A bilayer (“PIB-PU/E80A”) after 8 weeks (molecular weight) and 16 weeks (tensile, molecular weight) of the Oxidation Test.

Figure 7. Renderings of SEM images of the surface for ELASTHANE 80A polyether-polyurethane (Figure 7A) and ELAST-EON 5-135/ELASTHANE 80A bilayer (Figure 7C) tubing after 8 weeks of the oxidation test; Renderings of Optical images of cross-section of ELASTHANE 80A polyether-polyurethane (Figure 7B) and

ELAST-EON 5-135/ELASTHANE 80A bilayer (Figure 7D) tubing after 8 weeks of the Oxidation Test.

Figure 8. Molecular weight (weight-averaged) percentage of the ELASTHANE 80A polyether-polyurethane (“E80A control”; open squares) and ELAST-EON 5- 135/ELASTHANE 80A bilayer (“E5-E80A”; closed circles) up to 8 weeks of the oxidation test.

Figure 9. Mechanical (tensile) properties of the ELASTHANE 80A (“E80A”) and ELAST-EON 5-135/ELASTHANE 80A bilayer (“E5-E80A”) after 8 weeks of the oxidation test. Detailed Description

The present disclosure recognizes that there are tradeoffs in biostability and mechanical properties when designing a material for use in implantable medical devices. Typically, no one material can provide all the desirable properties, particularly no one soft 80A implantable polyurethane can provide all the desirable properties. Thus, the present disclosure uses a composite construct.

In one embodiment, there is provided an implantable medical device that includes: a substrate including a polyether-polyurethane; and a protective layer disposed directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer includes a more oxidatively-stable polyurethane. The protective material can be provided in one or more layers, if desired.

Significantly, the composite constructs (polyether-polyurethane substrate and protective layer) of the present disclosure retain the desirable hydrolytic stability and mechanical properties of soft polyether-polyurethanes, while improving their oxidative resistance. This is accomplished by employing a protective layer (typically, a surface layer) that includes a more oxidatively-stable polyurethane. The protective layer prevents the initiation and propagation of cracks from oxidation that could eventually lead to breaching.

This unique combination exploits the strengths of each material, while mitigating their weaknesses. While the more oxidatively-stable protective layer could itself degrade over the long-term, both chemically and mechanically (e.g., through hydrolysis), it is believed that it would provide protection from cell-mediated oxidation particularly during the initial critical period after implantation.

A balance of good adhesion between the polyurethane layers while still maintaining the soft flexible properties of the construct is important and allows for exploiting the strengths of each material, while mitigating their weaknesses. Thus, adhesion and stiffness of the material are important features to consider in selecting materials for use in the present disclosure.

The more oxidatively-stable polyurethanes suitable for use in the implantable medical devices of the present disclosure have a Shore A hardness of 50 to 95. More preferably, the more oxidatively-stable polyurethanes suitable for use in the implantable medical devices of the present disclosure have a Shore A hardness of 60 to 90, and even more preferably, a Shore A hardness of 70 to 85.

In some embodiments, the polyether-polyurethanes suitable for use in the implantable medical devices of the present disclosure have a Shore A hardness of 50 to 95. In some embodiments, the hardness of the polyether-polyurethane is outside this range, typically higher than this range (e.g., Shore D hardness of 55, which correlates to a Shore A hardness of 99). In some embodiments, the polyether- polyurethanes suitable for use in the implantable medical devices of the present disclosure have a Shore A hardness of 60 to 90, and even more preferably, a Shore A hardness of 70 to 85.

Ideally a construct where the more oxidatively-stable layer has similar stiffness to the polyether-polyurethane would be desirable. Although hardness (durometer) and stiffness (modulus) are not directly comparable mechanical properties, hardness is a suitable characterizing feature of stiffness for this purpose. Thus, for preferred embodiments, the more oxidatively-stable polyurethane has a hardness within 50 Shore A units, or within 40 Shore A units, or within 30 Shore A units, more preferably within 20 Shore A units, and even more preferably within 15 Shore A units, of that of the polyether-polyurethane.

A previous approach that has been utilized for improving the oxidative stability of implantable soft 80A polyether-polyurethanes has been to employ an outer layer of a harder, more oxidatively-stable, grade of polyether-polyurethane (e.g., 55D

PELLETHANE, which has a Shore D hardness level of 55 hardness level; a Shore A hardness of 100 is approximately equal to a Shore D hardness level of 58, and a Shore A hardness of 95 is approximately equal to a Shore D hardness of 46). The harder grade polyether-polyurethane contains less of the polyether soft segment, which is known to be more susceptible to oxidation. The compatibility and

consequently adhesion between these layers is excellent as they are chemically similar and can readily mix and diffuse into one another (e.g., through chain entanglement) under typical processing conditions (e.g., thermal co-extrusion, etc.). However, the higher modulus of the more oxidatively-stable coating could create a bigger stress riser at the crack tip, if a crack is initiated on the surface. This large stress riser could drive the crack into the softer under layer and expose the more susceptible material to the oxidative environment. Another consideration with this approach is the increased stiffness of the construct, which may be less desirable for applications where significant flexibility is required, for example in the vasculature (tortuous path), in the heart (flexing), and in tissues that may be sensitive to stiffer materials (e.g., brain). With that said, adding a soft oxidatively stable layer on top of the harder grade polyether-polyurethane could prevent or decrease surface crack formation and protect the underlying harder material. In the present disclosure, in addition to“matching” the softness of the polyether-polyurethanes and more oxidatively-stable polyurethanes, the polyether- polyurethanes and the more oxidatively-stable polyurethanes suitable are selected to have suitable adhesion to each other. Adhesion between layers is particularly important as the construct will be exposed to mechanical loading (flexing, etc.), which if sufficient, could overcome bond strength resulting in delamination. Generally, the more-oxidatively stable material being from the polyurethane family facilitates good adhesion to the underlying polyether-polyurethane due to similar solubility of components (which promotes mixing, chain entanglement, etc.). Suitable adhesion can be demonstrated by performing an oxidative stress test and evaluating for evidence of delamination between layers (e.g., voids), as demonstrated in the

Examples Section.

For the purposes of defining a more oxidatively-stable polyurethane compared to the selected polyether-polyurethane, relative performance can be used with respect to retention of physical, mechanical, and chemical properties. This can be done using an in vitro accelerated oxidation test commonly known as the“Stokes Test” developed by Ken Stokes and Jim Anderson (Schubert et al., J. Biomed. Mater. Res., 34, 519-530 (1997), and Zhao et al., J. Biomed. Mate.r Res., 29, 467-475 (1995)).

It is recognized that there are several variants of this test, however, for the purposes of comparison, the test that is preferred and which is used to compare materials of this invention (“the oxidation test”) is hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M) solution at 37^oC for up to 16 weeks (without glass wool). The samples are monitored periodically (typically, at 2, 4, and 8 weeks) and at the end of this 16-week period for changes compared with control unexposed samples (surface morphology, bulk mechanical properties, chemical, etc.). The cobalt chloride catalyzes the formation of reactive oxygen-based free radicals from hydrogen peroxide (hydroxyl (OH), superoxide (O - 2), peroxyl (HOO), etc.) which can readily attack the polyether soft segment of polyether-polyurethanes. Many of these reactive oxygen-based free radicals are believed to be similar to those secreted by inflammatory cells

(macrophages, foreign body giant cells, etc.) in vivo. The test that has been shown to accelerate the effects of oxidation in vivo with similar changes observed chemically and in surface morphology (Christenson et al., J. Biomed. Mater. Res., 70A, 245-255 (2004)).

Characterization methods that could be used to asses changes due to oxidation include: Scanning Electron Microscopy (SEM) to observe changes in morphology, such as pitting and micro-cracks (crazing), as shown in Examples 1 and 4, or macroscopic visual changes (cracking, breaching), mechanical changes (tensile properties as shown in Example 5), chemical changes via spectroscopic methods including FT-IR (Fourier Transform Infrared) spectroscopy showing loss of ether soft segment content (reduction in peak at 1105 cm^-1 relative to urethane ether at 1075 cm^-1), and/or increase in the ratio of hydrogen bonded carbonyl (1700 cm^-1) relative to free carbonyl (1730 cm^-1), changes in the molecular weight distribution from oxidation processes resulting in chain scission and/or crosslinking. It is further recognized that there are other chemical (Nuclear Magnetic Resonance– NMR, etc.) and mechanical (Dynamic Mechanical Analysis– DMA, etc.) techniques that could be utilized to assess changes due to oxidation.

The polyether-polyurethanes are well-known materials in the implantable medical device industry. A typical polyether-polyurethane is made from methylene diphenyl diisocyanate (MDI), butane diol (BDO), and polyether macrodiols such as poly(tetramethylene oxide) (PTMO). Examples of these include those available under the tradenames ELASTHANE 80A & 55D from DSM Biomedical, and

PELLETHANE 236380A & 236355D from Lubrizol. ELASTHANE

80A/PELLETHANE 80A polyether-polyurethanes possess 60% by weight of polyether content. Typically, the polyether forms the soft segment of the polymer (i.e., a portion of the polyurethane and has a glass transition temperature (Tg) that is below body temperature).

The more oxidatively-stable polyurethanes are polyurethane multi-block polymers that show increased stability, compared with polyether-polyurethanes, to chemical and physical changes when exposed to an oxidative environment.

Oxidative stability can be measured by physical changes including surface pitting or cracking, mechanical property changes, and chemical changes including molecular weight changes, changes in chemical composition, etc. Preferably, the more oxidatively-stable polyurethane is selected from the group of a PDMS-polyurethane, a polycarbonate-polyurethane, a polyisobutylene- polyurethane, or combinations thereof. In this context,“combinations thereof” refer to polymers that contain mixed functionality (e.g. silicone carbonate-polyurethane), or blends of such polymers.

Examples of PDMS-polyurethanes include those available under the

tradenames ELAST-EON (e.g., ELAST-EON 2A, ELAST-EON 5-130, ELAST-EON 5- 325, ELAST-EON 5-135, ELAST-EON 2-852, ELAST-EON 2-860, and ELAST-EON 2-862), and EC-SIL (e.g., EC-SIL 70A, EC-SIL 75A, EC-SIL 90A) from Aortech.

Examples of PDMS-polyurethanes include those available under the tradenames PURSIL (such as PURSIL with 10% to 40% PDMS content) and CARBOSIL (such as CARBOSIL with 5% to 20% PDMS content) from DSM Biomedical. In such polymers, the PDMS content typically replaces a portion or all of a polyether or polycarbonate soft segment.

Examples of polycarbonate-polyurethanes include those available under the tradenames BIONATE (e.g., BIONATE PCU 80A, BIONATE PCU 90A, and

BIONATE II) from DSM Biomedical, and QUADRATHANE from Biomerics.

Examples of polyisobutylene-polyurethanes can be prepared according to the procedures described in U.S. Pub. No. 2010/0130696 A1 to Kennedy et al, and U.S. Pub.No. 2010/0179298 A1 to Faust et al.

The more oxidatively-stable polyurethane can be present as a single

component in the protective layer or it can be blended with other polymers in the protective layer. Typically, the protective layer includes an amount of the more oxidatively-stable polyurethane such that the overall polyether content of the protective layer is less than 60% by weight (wt-%) in the blend.

A preferred polymer construct of the present disclosure includes a soft 80A polyether-polyurethane (e.g., such as those available under the tradenames

ELASTHANE 80A or PELLETHANE 80A) as the substrate layer and a soft protective layer that includes a more oxidatively-stable polyurethane. The more oxidatively- stable protective layer provides oxidative protection and prevents or significantly reduces the formation of surface cracks that could propagate and lead to breaching of the underlying substrate layer. The substrate material, for example, ELASTHANE 80A polyether-polyurethane, would comprise the main overall thickness of the construct, thus providing mechanical strength and support for the well-adhered protective layer. The protective layer preferably includes one or more oxidatively- stable polyurethanes with hardness ranging from Shore 50A to 95A. Materials for this protective (typically, surface) layer preferably include a PDMS-polyurethane, a polycarbonate-polyurethane, and a polyisobutylene-polyurethane. These

polyurethanes have good adhesion to the substrate polyether-polyurethane due to their similar chemistry.

In certain embodiments, preferred constructs will maintain the same

softness/stiffness as polyether-polyurethane 80A, and exhibit good biocompatibility.

Also, in certain embodiments, preferred constructs overcome the problems facing a single polyurethane material, for example, loss of mechanical strength due to oxidation of a polyether-polyurethane, loss of mechanical strength due to hydrolysis of a PDMS-polyurethane/polycarbonate polyurethane, and/or high material cost and lower mechanical strength of a polyisobutylene-polyurethane.

It is further recognized that the nature of the oxidative process could be confined to a few microns of the surface. Thus, according to the present disclosure, a thin layer of the more oxidatively-stable polyurethane can be used. Typically, a layer at least 2 microns thick, and often up to 50 microns thick, or up to 10 microns this, is used to significantly reduce the oxidative attack and completely inhibit the surface crack formation, which significantly reduces the degradation of the

mechanical properties of the polyether-polyurethane substrate. It is recognized that the thickness of the more oxidatively-stable polyurethane layer may be limited by the process used for preparing the construct.

The more oxidatively-stable surface layer is preferably less than 50% of the overall thickness of the construct, and more preferably 25% or less of the overall thickness, but typically greater than 1 % of the overall thickness. In this way, the main bulk mechanical properties of the construct are typically provided by the base polyether-polyurethane substrate.

In the past, various surface treatments to polyether-polyurethanes have been tried to slow or prevent its biodegradation, such as absorbing albumin, grafting of poly (2-hydroxyethyl-methacrylate), and absorbing silicone to the surface. However, none of these approaches have been shown to be effective. Surface absorbed PDMS on polyether polyurethane fiber showed reduced cracking but it also showed delamination, which is not desired (“The use of silicone/polyurethane graft polymers as a means of eliminating surface cracking of polyurethane prostheses.” L. Pinchuk et al., J. Biomater Appl., 3:260 (1988)). This led to the development of covalent bonding of PDMS to the polyurethane surface for better bonding, however, this process has not been widely adopted. It is suspected that difficulties in processing, surface properties, and performance were encountered (“Polyurethane/silicone composites for the long term implant in the human body.” L. Pinchuk et al., ANTEC, P1802 (1991 )).

The present disclosure provides processing advantages as well as

performance advantages. Suitable methods of making constructs of the present disclosure include: 1 ) thermal processes including extrusion processes such as co- extrusion, over-extrusion, lamination, etc.; 2) solvent or dispersion coating processes such as dip coating or spray coating, etc.; 3) powder coating methods; and 4) plasma deposition processes.

Co-extrusion is the process of extruding two different materials at the same time through a single die with two or more orifices arranged so that the extrudates merge and weld together into a laminar structure before chilling. Each material is fed to the die from a separate extruder, but the orifices may be arranged so that each extruder supplies two or more plies of the same material. In this case, polyether- polyurethane and an oxidatively-stable polyurethane (PDMS-polyurethane, polyisobutylene-polyurethane, polycarbonate-polyurethane, etc.) are put into different extruders and co-extruded with polyether-polyurethane as the base substrate layer and the oxidatively-stable polyurethane as the surface layer. In a preferred tubing form, polyether-polyurethane will be inside and the more oxidatively-stable

polyurethane will be on the outside (tubing form). Over-extrusion involves extruding a second material over a first material which was previously processed into the desired form. Lamination involves two materials being mated together using heat and typically pressure. Other thermal processes for laminating (fusing) layers that could be envisaged include heat shrink tubing, etc. Dip coating, spray coating, powder coating, or plasma deposition can also be used to form a layer including a more oxidatively-stable polyurethane on the polyether-polyurethane substrate. For example, extruded polyether-polyurethane tubing can be coated with a PDMS-polyurethane by passing the tubing through a solution or dispersion of the PDMS-polyurethane. Typical solvents for preparing coating solutions of the more oxidatively-stable polyurethanes include THF

(tetrahydrofuran), DMF (dimethyl formamide), NMP (N-methyl pyrrollidinone), DMAC (dimethyl acetamide), etc. Solutions of the more oxidatively-stable polyurethanes can be prepared prior to coating or obtained already pre-made, for example, ELAST- EON 2A (AorTech Inc.) is available from the manufacturer in solution grade and is soluble in both THF and DMAc. The more oxidatively-stable polyurethane can also be applied on the polyether-polyurethane by spray coating methods. A solution or dispersion of the selected more oxidatively-stable polyurethane can be sprayed over tubing of polyether-polyurethane, for example.

The polymer constructs (the polyether-polyurethane substrate and the more oxidatively-stable polyurethane protective layer) can be used as insulation disposed on an electrical lead, such as a cardiac pacing or defibrillator lead, etc. They can also be used as lead insulation for neurostimulation leads (e.g., deep brain stimulation leads, spinal cord stimulation leads), or other electrical stimulation leads. See, for example, Figure 1 , which shows schematics of (A) a cardiac lead and (B) a deep brain stimulation (DBS) lead (top view).

Figure 1 A is a schematic view of an exemplary implantable medical device (IMD) 10 including atrial lead 12 and ventricular lead 14 implanted in heart 16 (e.g., as described in greater detail in U.S. Pat. No. 7,780,607). IMD 10 may be a pacemaker, defibrillator, cardioverter, pacemaker/cardioverter/defibrillator (PCD), heart function monitor having pacing capabilities, or other implantable device that includes the capability of providing therapy to heart 16. IMD 10 includes connector module or header 18 and housing 20. Atrial lead 12 and ventricular lead 14 extend from connector module 18 into the right atrium RA and right ventricle RV,

respectively, of heart 16. Proximal ends of atrial lead 12 and ventricular lead 14 are connected at header 18 to sensing, signal processing, and therapy delivery circuitry (not shown) within housing 20. Atrial lead 12 and ventricular lead 14 enter right atrium RA through superior vena cava 24. Atrial lead 12 is a J-shaped bipolar lead including tip electrode 30 and ring electrode 32 at its distal end, while ventricular lead 14 is an elongated bipolar lead including tip electrode 34 and ring electrode 36 at its distal end. While bipolar leads 12 and 14 are disclosed, unipolar leads can alternatively be implanted in the same anatomic relation to the heart chambers.

When heart 16 contracts, atrial lead 12 and ventricular lead 14 are deflected. The atrial contraction causes bending or deformation of atrial lead 12 along bending portion 40, while the ventricular contraction causes bending or deformation of ventricular lead 14 along bending portion 42. The magnitude of the deflection along bending portions 40 and 42 depends on the radial stiffness of atrial lead 12 and ventricular lead 14, respectively, and on the muscle contraction forces of heart 16. In addition, the magnitude of the deflection depends on the initial bending forces caused by the specific implantation position. For instance, atrial lead 12 implanted on the anterior atrial wall (as shown in Figure 1 A) has a larger J-shape radius than a lead implanted in the atrial appendage. Atrial lead 12 and ventricular lead 14 are strongly mechanically coupled to the heart muscle, especially in the chronic phase of cardiac pacing when fibrotic tissue anchors the lead tips to the endocardium. The polymer constructs (the polyether-polyurethane substrate and the more oxidatively- stable polyurethane protective layer) can be used as insulation disposed on atrial lead 12 and ventricular lead 14.

Figure 1 B illustrates a DBS system implanted in a patient 40 that includes at least one neurostimulator, at least one extension, and at least one stimulation lead containing electrodes (e.g., as described in greater detail in EP1740260 B1 ). As can be seen, each neurostimulator 42 is implanted in the pectoral region of the patient. Extensions 44 are deployed up through the patient's neck, and leads 46 are implanted in the patient's brain as shown at 48. As can be seen, each of the leads 46 is connected to its respective extension 44 just above the ear on both sides of patient 40. The polymer constructs (the polyether-polyurethane substrate and the more oxidatively-stable polyurethane protective layer) of the present disclosure can be used as insulation disposed on leads 46.

The polymer constructs (the polyether-polyurethane substrate and the more oxidatively-stable polyurethane protective layer) can be used as catheter tubing, such as drug infusion catheter tubing (i.e., drug delivery catheter tubing), hemodialysis catheter tubing, peritoneal catheter tubing, etc.

Exemplary embodiments of some drug delivery systems for infusing drugs are depicted in Figures 2A and 2B (e.g., as described in greater detail in U.S. Pat. No. 8,043,281 ). Figure 2A is a schematic diagram of a drug delivery system for infusing drug to the brain, and Figure 2B is a schematic diagram of a drug delivery system for infusing drug to the spinal region. The drug delivery systems depicted in Figures 2A and 2B includes a drug infusion pump assembly 10A/10B and catheter 20A/20B having a proximal end 22A/22B attached to the pump assembly and distal end 24A/24B implanted within the patient. The distal end 24A is implanted within the brain 30A of the patient, while the distal end 24B is implanted within the spinal column 30B of the patient.

The polyether-polyurethane substrate typically forms the tubing body (e.g., of catheter 20A/20B in Figures 2A and 2B) having an inner surface and an outer surface, and the more oxidatively-stable polyurethane protective layer can be directly disposed on the outer surface of the tubing body (as shown in Figure 3, illustration on the left), or on the inner surface of the tubing body, or both (as shown in Figure 3, illustration on the right).

Thus, a further embodiment of this disclosure is that the more oxidatively- stable polyurethane protective layer could also be applied on the inner surface of the construct (tubing, etc.). This may be useful in preventing oxidative processes from occurring such as MIO (from exposed metal lead conductors) or long-term exposure to drug formulations (e.g., drug infusion catheters). Figure 3 (the illustration on the right) shows the coating applied to both the inner (ID) and outer (OD) diameters.

The polymer constructs (the polyether-polyurethane substrate and the more oxidatively-stable polyurethane protective layer) can be used as layers on a medical device, such that the polyether-polyurethane substrate is in the form of a layer of material.

The implantable medical device can be a pacemaker, cardiac defibrillator, neuromodulation stimulator (e.g., deep brain stimulator, spinal cord stimulator, etc.), drug infusion catheter, heart valve, stent, an orthopedic product (e.g., for cranial repair, or a spine correction rod, plate, or screw), hydrocephalus shunt catheter, or an artificial spinal disc. Exemplary Embodiments

Emobiment 1 is an implantable medical device comprising: a substrate comprising a polyether-polyurethane; and a protective layer disposed directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer comprises a more oxidatively-stable polyurethane; wherein the more

oxidatively-stable polyurethane has a Shore A hardness of 50 to 95.

Embodiment 2 is the implantable medical device of embodiment 1 wherein the more oxidatively-stable polyurethane has a hardness within 50 Shore A units of the hardness of the polyether-polyurethane.

Embodiment 3 is the implantable medical device of embodiment 1 or 2 wherein the more oxidatively-stable polyurethane is selected from the group of a PDMS-polyurethane, a polycarbonate-polyurethane, a polyisobutylene-polyurethane, or combinations thereof.

Embodiment 4 is the implantable medical device of any of embodiments 1 through 3 wherein the protective layer comprises a blend.

Embodiment 5 is the implantable medical device of embodiment 4 wherein the blend comprises an amount of the more oxidatively-stable polyurethane such that the polyether content is less than 60 wt-% in the blend.

Embodiment 6 is the implantable medical device of any of embodiments 1 through 5 further comprising an electrical lead, wherein the polyether-polyurethane substrate and protective layer form insulation disposed on the electrical lead.

Embodiment 7 is the implantable medical device of embodiment 6 wherein the electrical lead is a neurostimulation lead, a cardiac pacing lead, or a defibrillator lead.

Embodiment 8 is the implantable medical device of any of embodiments 1 through 5 wherein the polyether-polyurethane substrate and protective layer form catheter tubing.

Embodiment 9 is the implantable medical device of embodiment 8 wherein the catheter tubing is drug infusion catheter tubing. Embodiment 10 is the implantable medical device of embodiment 8 or 9 wherein the polyether-polyurethane substrate forms the tubing body having an inner surface and an outer surface, and the protective layer is directly disposed on the outer surface of the tubing body.

Embodiment 11 is the implantable medical device of any of embodiments 8 through 10 wherein the polyether-polyurethane substrate forms the tubing body having an inner surface and an outer surface, wherein the protective layer is directly disposed on the inner surface of the tubing body.

Embodiment 12 is the implantable medical device of any of embodiments 1 through 5 wherein the polyether-polyurethane substrate forms a layer on a medical device.

Embodiment 13 is the implantable medical device of any of embodiments 1 through 12 wherein the polyether-polyurethane and the protective layer are formed by a thermal process.

Embodiment 14 is the implantable medical device of any of embodiments 1 through 12 wherein the protective layer is applied to the polyether-polyurethane substrate by solution coating, dispersion coating, plasma deposition, or powder coating.

Embodiment 15 is the implantable medical device of any of embodiments 1 through 5 or embodiments 12 through 14 which is selected from the group of a pacemaker, cardiac defibrillator, neuromodulation stimulator, drug infusion catheter, heart valve, stent, an orthopedic product, hydrocephalus shunt catheter, and an artificial spinal disc.

Embodiment 16 is the implantable medical device of any of embodiments 1 through 15 wherein wherein the polyether-polyurethane and the more oxidatively- stable polyurethane each have a Shore A hardness of 50 to 95.

Embodiment 17 is an implantable medical device comprising: a polyether- polyurethane layer; and a protective layer disposed directly on at least one surface of the polyether-polyurethane layer, wherein the protective layer comprises a PDMS- polyurethane.

Embodiment 18 is the implantable medical device of embodiment 17 wherein the polyether-polyurethane layer and the protective layer are co-extruded layers. Embodiment 19 is a method of improving the oxidative stability of an implantable medical device, the method comprising: providing a substrate comprising a polyether-polyurethane; and applying a protective layer directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer comprises a more oxidatively-stable polyurethane.

Embodiment 20 is the method of embodiment 19 wherein the polyether- polyurethane layer and the protective layer are co-extruded.

Embodiment 21 is the implantable medical device of any of embodiments 1 through 18 or the method of embodiment 19 or 20 wherein the protective layer is at least 2 microns thick.

Examples

Objects and advantages of this disclosure are further illustrated by the following examples, but the particular materials and amounts thereof recited in these examples, as well as other conditions and details, should not be construed to unduly limit this disclosure.

The following examples demonstrate that different polyurethane materials can be combined into a multi-layer construct to improve biostability while retaining the desirable mechanical and processing properties of polyurethanes. Materials

A thermoplastic polyether-polyurethane elastomer available under the tradename ELASTHANE 80A and a thermoplastic polycarbonate polyurethane available under the tradename BIONATE were purchased from DSM Biomedical, and poly(dimethyl siloxane)-polyurethanes available under the tradenames ELAST-EON 2A (E2A) and ELAST-EON 5-135 were obtained from AorTech.

Polyisobutylene-polyurethane 75A was custom synthesized following the general method outlined in the publication by Cozzens et al. (J. Biomed. Mater. Res. Part A: 95A:774-782 (2010)) using a PIB-diol with Mn of 3.5 kDa, which was synthesized by the method described in the publication by Unmadisetty et al. (Journal of Polymer Science: Part A: Polymer Chemistry, Vol. 46, 4236-4242 (2008)). The product included 67 wt-% soft segment (80% polyisobutylene (PIB) + 20%

polytetramethylene oxide (PTMO)) and 33 wt-% hard segment (81 % methylene diphenyl diisocyanate (MDI) + 19% 1 ,4-butanediol (BDO)). Example 1.

Soft elastomers with hardness in the range of Shore 80–90A were molded into flat sheets and immersed in an oxidative hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M) solution at 37ºC (i.e.,“The Oxidation Test”). After two weeks, the samples were removed, rinsed and imaged by SEM. Significant surface pitting which is indicative of oxidation was observed for ELASTHANE 80A polyether-polyurethane, whereas very little or no significant pitting was observed for ELAST-EON 2A poly(dimethyl siloxane)-polyurethane, ELAST-EON 5-135 poly(dimethyl siloxane)-polyurethane, BIONATE II, and polyisobutylene-polyurethane, indicating improved oxidation stability compared to ELASTHANE 80A polyether-polyurethane.

Additional samples were immersed for 8 & 16 weeks in the same peroxide solution. Figure 4 shows the renderings of SEM images for PELLETHANE 80A (Figure 4A, Figure 4B) and ELAST-EON 2A (Figure 4C, Figure 4D) after exposure to hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M) for an 8-week period (Figure 4A and Figure 4C, 400X for both) and a 16-week period (40X for PELLETHANE 80A (Figure 4B) and 400X for ELAST-EON 2A (Figure 4D)). The increased duration of peroxide exposure showed a progression from pitting after 2 weeks, to cracking after 8 weeks, which appeared to progress further after 16 weeks for polyether polyurethanes (PELLETHANE 80AE, ELASTHANE 80A). Cracking is evidence of further oxidation. More specifically, PELLETHANE 80A and ELASTHANE 80A polyether-polyurethanes exhibited cracking after 8-wk exposure, and progression of crack growth was evident after 16-wk exposure. For ELAST-EON 2A poly(dimethyl siloxane)-polyurethane and polyisobutylene-polyurethane no cracks were observed after 16 weeks with only minor pitting observed.

Also, after 16 weeks in the test solution, ELAST-EON 2A poly(dimethyl siloxane)-polyurethane, ELAST-EON 5-135 poly(dimethyl siloxane)-polyurethane, BIONATE II polycarbonate-polyurethane, and polyisobutylene-polyurethane showed significantly less molecular weight decrease than ELASTHANE 80A polyether- polyurethane. Example 2.

ELAST-EON 2A poly(dimethyl siloxane)-polyurethane solution (100 mg/mL) in THF was used to coat a ELASTHANE 80A polyether-polyurethane (sheet of thickness 0.65 mm) The coating was prepared via a Meyer-rod coater. The thickness of the ELAST-EON 2A poly(dimethyl siloxane)-polyurethane layer was estimated at approximately 4 μm which was confirmed by Raman confocal microscopy. Raman confocal microscopy showed a uniform thin layer of ELAST-EON 2A poly(dimethyl siloxane)-polyurethane was coated on the ELASTHANE 80A polyether-polyurethane. Example 3.

ELASTHANE 80A polyether-polyurethane was molded into flat sheets with thickness around 0.6 mm. Polyisobutylene-polyurethane 80A, ELAST-EON 2A poly(dimethyl siloxane)-polyurethane, and SIBS were prepared with a concentration of 10 wt-%, and coated on the ELASTHANE 80A polyether-polyurethane similarly as described in Example 2. After drying, samples were immersed in hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M) solution at 37ºC (i.e.,“The Oxidation Test” set forth in

Example 1 ). After 4 weeks, delamination was found for the styrene-isobutylene- styrene (SIBS) coated sample (small bubbles visible at the interface between the layers).

Styrene-isobutylene-styrene copolymers (SIBS) with polyurethanes have been proposed to be used to construct implantable medical device insulators, although no examples were given (U.S. Pub. No. 2010/0076538 A1 ). However, being chemically dissimilar, adhesion between the SIBS olefinic and urethane layers would not be expected to be robust. In fact, this example demonstrates that SIBS has poor adhesion to polyurethane and is prone to delamination when used as a surface coating on the polyether-polyurethane. Example 4.

The ELASTHANE 80A polyether-polyurethane sample coated with ELAST- EON 2A poly(dimethyl siloxane)-polyurethane (E2A) from Example 2 was evaluated in hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M) at 37ºC to evaluate the protection effect of ELAST-EON 2A poly(dimethyl siloxane)-polyurethane. Samples immersed in the solution for 2 and 4 weeks were taken out and analyzed by SEM. After 4 weeks no significant changes in surface morphology were observed for the E2A- coated side, whereas the uncoated ELASTHANE 80A side showed significant pitting. These observations show that E2A provided significant protection to the

ELASTHANE 80A surface from oxidation. Example 5.

ELASTHANE 80A polyether-polyurethane (referred to herein as“E80A” or “PEU80A”) sheets were coated with ELAST-EON 2A poly(dimethyl siloxane)- polyurethane (E2A) on both sides, and immersed in the hydrogen peroxide (20 wt- %)/CoCl₂ (0.1 M) solution at 37ºC for 16 weeks (the“oxidation test”).

Both bare PEU80A and E2A-coated PEU80A were evaluated after peroxide exposure.

After 16-week exposure of hydrogen peroxide (20 wt-%)/CoCl₂ (0.1 M), the bare PEU80A had cracks while E2A-coated PEU80A did not exhibit cracks. Confocal Raman microscopy showed that the E2A coating was approximately 4 microns thick Tensile evaluation was performed via ASTM D638-10 using an MTS Sintech 1 /D test frame with extensometer. After exposure to peroxide solution, samples were cut from the coated (both sides) and bare control sheets (overall thickness

approximately 0.025 inch) using an ASTM D638-10 Type V die and pulled at a crosshead speed of 5 inches/min (n = 10). The samples were tested in a hydrated condition at ambient temperature. Tensile evaluation showed E2A-coated PEU80A (“Coated E80A”) has significantly reduced degradation in mechanical properties compared with bare PEU80A (“Bare E80A”), as shown in Figure 5. Example 6.

Polyether-polyurethane tubing is coated with one the candidate materials by dip or spray coating in solvents such as THF (tetrahydrofuran), DMF (dimethyl formamide), NMP (N-methyl pyrrollidinone, and DMAC (dimethyl acetamide). The thickness of the coating is controlled by the concentration of the polymer solution. Both the inside and outside of the tubing can be coated or just outside of tubing is coated as illustrated in Figure 3. Example 7.

A 10% solution of polyisobutylene-polyurethane (PIB-PU) in THF was prepared as the coating solution and both sides of ELASTHANE 80A sheet sample was coated with this PIB-PU solution. Wet thickness of the coating was 50 μm and the coating was quickly dried in an 80ºC oven. The dried PIB-PU coating layer was estimated to be around 4-5 μm thick. Both ELASTHANE 80A sheet with and without PIB-PU coating (control) were tested for oxidation stability using the Oxidation Test described in Example 1. After 8-week and 16-week time points, samples were taken out for molecular weight measurements and tensile property evaluations. The tensile data and weight averaged molecular data are shown in Figure 6. The results presented in Figure 6 clearly show that a thin coating of PIB-PU significantly protected the underlying ELASTHANE 80A base layer from oxidation (“PIB- PU/E80A”), thereby maintaining significant mechanical properties over longer time periods compared with ELASTHANE 80A control without coating (“E80A”). Example 8.

One example of bilayer tubing was co-extruded with ELASTHANE 80A polyether-polyurethane as the inside base layer targeting thickness of 100 microns and ELAST-EON 5-135 as the outer layer targeting thickness of 25 microns. The tubing was co-extruded by feeding the melted resins into two separate extruders. The extruded tubing was examined by looking at the cross-section using optical microscope and SEM. Targeted thickness was achieved by optical microscope and SEM with the two layers showing good adhesion between them (no signs of delamination at the interface). Example 9.

In another case, the bilayer is co-extruded with ELASTHANE 55D polyether- polyurethane as the inside base layer and PURSIL 35 as the outer layer. The tubing is co-extruded using a similar method as in Example 8. The extruded tubing is examined by looking at the cross-section using optical microscope and SEM.

Targeted thickness is achieved by optical microscope and SEM. Example 10.

The co-extruded tubing from Example 8 was tested in the oxidation solution using ELASTHANE 80A polyether-polyurethane tubing as a control using the

Oxidation Test described in Example 1. Figure 7 shows renderings of SEM images of the surface for ELASTHANE 80A (Figure 7A) and ELAST-EON 5- 135/ELASTHANE 80A bilayer (Figure 7C) tubing after 8 weeks of the oxidation test, and renderings of optical images of cross-section of ELASTHANE 80A (Figure 7B) and ELAST-EON 5-135/ELASTHANE 80A bilayer (Figure 7D) tubing after 8 weeks of the Oxidation Test. ELASTHANE 80A polyether-polyurethane tubing showed surface cracks after 8 weeks as shown in Figure 7 (A and B), while the ELAST-EON 5-135/ELASTHANE 80A bilayer tubing prevented crack formation and maintained good adhesion between the ELAST-EON 5-135 surface and ELASTHANE 80A base layers as shown in Figure 7 (C and D).

The molecular weights of the tubing samples were tracked with time using GPC as shown in Figure 8. After 8 weeks, ELAST-EON 5-135/ELASTHANE 80A bilayer tubing (“E5-E80A”) showed a modest 15% decrease of the weight averaged molecular weight, whereas the ELASTHANE 80A polyether-polyurethane (“E80A control”) showed a significant 75% reduction.

The mechanical (tensile) properties of the tubing samples were also tested after 8 weeks of the oxidation test. The ELAST-EON 5-135/ELASTHANE 80A (“E5- E80A”) bilayer tubing maintained more than 4000 PSI ultimate tensile strength (UTS) and 400% elongation at break, however, the ELASTHANE 80A polyether- polyurethane (“E80A”) tubing had less than 100 PSI and 50% elongation at break, as shown in the Figure 9. The complete disclosures of the patents, patent documents, and publications cited herein are incorporated by reference in their entirety as if each were individually incorporated. Various modifications and alterations to this disclosure will become apparent to those skilled in the art without departing from the scope and spirit of this disclosure. It should be understood that this disclosure is not intended to be unduly limited by the illustrative embodiments and examples set forth herein and that such examples and embodiments are presented by way of example only with the scope of the disclosure intended to be limited only by the claims set forth herein as follows.

Claims

WHAT IS CLAIMED IS:

1. An implantable medical device comprising:

a substrate comprising a polyether-polyurethane; and

a protective layer disposed directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer comprises a more oxidatively-stable polyurethane;

wherein the more oxidatively-stable polyurethane has a Shore A hardness of 50 to 95.

2. The implantable medical device of claim 1 wherein the more oxidatively-stable polyurethane has a hardness within 50 Shore A units of the hardness of the polyether-polyurethane.

3. The implantable medical device of claim 1 wherein the more oxidatively-stable polyurethane is selected from the group of a PDMS-polyurethane, a

polycarbonate-polyurethane, a polyisobutylene-polyurethane, or combinations thereof.

4. The implantable medical device of claim 1 wherein the protective layer

comprises a blend.

5. The implantable medical device of claim 4 wherein the blend comprises an amount of the more oxidatively-stable polyurethane such that the polyether content is less than 60 wt-% in the blend.

6. The implantable medical device of claim 1 further comprising an electrical lead, wherein the polyether-polyurethane substrate and protective layer form insulation disposed on the electrical lead.

7. The implantable medical device of claim 6 wherein the electrical lead is a

neurostimulation lead, a cardiac pacing lead, or a defibrillator lead.

8. The implantable medical device of claim 1 wherein the polyether-polyurethane substrate and protective layer form catheter tubing.

9. The implantable medical device of claim 8 wherein the catheter tubing is drug infusion catheter tubing.

10. The implantable medical device of claim 8 wherein the polyether-polyurethane substrate forms the tubing body having an inner surface and an outer surface, and the protective layer is directly disposed on the outer surface of the tubing body.

11. The implantable medical device of claim 8 wherein the polyether-polyurethane substrate forms the tubing body having an inner surface and an outer surface, wherein the protective layer is directly disposed on the inner surface of the tubing body.

12. The implantable medical device of claim 1 wherein the polyether-polyurethane substrate forms a layer on a medical device.

13. The implantable medical device of claim 1 wherein the polyether-polyurethane and the protective layer are formed by a thermal process.

14. The implantable medical device of claim 1 wherein the protective layer is

applied to the polyether-polyurethane substrate by solution coating, dispersion coating, plasma deposition, or powder coating.

15. The implantable medical device of claim 1 which is selected from the group of a pacemaker, cardiac defibrillator, neuromodulation stimulator, drug infusion catheter, heart valve, stent, an orthopedic product, hydrocephalus shunt catheter, and an artificial spinal disc.

16. The implantable medical device of claim 1 wherein wherein the polyether- polyurethane and the more oxidatively-stable polyurethane each have a Shore A hardness of 50 to 95.

17. An implantable medical device comprising:

a polyether-polyurethane layer; and

a protective layer disposed directly on at least one surface of the polyether-polyurethane layer, wherein the protective layer comprises a PDMS- polyurethane.

18. The implantable medical device of claim 17 wherein the polyether- polyurethane layer and the protective layer are co-extruded layers.

19. A method of improving the oxidative stability of an implantable medical device, the method comprising:

providing a substrate comprising a polyether-polyurethane; and applying a protective layer directly on at least one surface of the polyether-polyurethane substrate, wherein the protective layer comprises a more oxidatively-stable polyurethane.

20. The method of claim 19 wherein the polyether-polyurethane layer and the

protective layer are co-extruded.

21. The method of claim 19 wherein the protective layer is at least 2 microns thick.