WO2008140376A1

WO2008140376A1 - Tantalum electrode

Info

Publication number: WO2008140376A1
Application number: PCT/SE2008/000311
Authority: WO
Inventors: Andreas ÖRNBERG
Original assignee: St. Jude Medical Ab
Priority date: 2007-05-14
Filing date: 2008-05-07
Publication date: 2008-11-20

Abstract

The invention relates to implantable electrodes (28, 30, 38) and in particular such electrodes (28, 30, 38) for usage in connection with medical electrical leads (10). The electrodes (28, 30, 38) of the invention comprises an electrode core (60) of tantalum or a biocompatible tantalum alloy. At least a portion of the outer surface (65) of the core (60) is coated with a conducting layer (70) that inhibits the growth of a resistive oxide on the surface potion. The conducting layer (70) is made of a conductive nitride, oxide, carbide, carbonitride or bromide.

Description

TANTALUM ELECTRODE

TECHNICAL FIELD

The present invention generally relates to implantable electrodes, and in particular to tantalum electrodes for usage in connection with medical electrical leads.

BACKGROUND

Various types of body-implantable leads are known and used in the medical field. For example, implantable medical devices, IMDs, such as pacemakers, cardiac defibrillators and cardioverters are, in operation, connected to implantable leads for sensing cardiac function and other diagnostic parameters and delivering stimulation pulses. For example, endocardial leads are attached at their proximal end to an IMD and at their distal end to the endocardium of a cardiac chamber.

The prior art implantable electrical leads typically use platinum or a platinum-iridium alloy, often no more than 30 % iridium, as electrode material. Although, platinum has several advantageous properties making it the main electrode material of today, it has a big disadvantage in terms of cost. Platinum is a very expensive metal, leading to high overall costs for implantable leads having platinum or platinum/ iridium electrodes.

Documents [5, 6, 8, 17, 18] discuss the usage of tantalum in connection with capacitor stimulation electrodes that can be used in different implantable medical devices, including pacemakers. The tantalum electrodes are covered with a dielectric, non-conducting tantalum pentoxide (Ta2θs) surface layer. However, such an electrode suffers from hydrogen embrittlement if used as a cathode in a corrosion cell.

SUMMARY

The present invention overcomes these and other drawbacks of the prior art arrangements. It is a general object of the present invention to provide an implantable tantalum electrode.

It is another object of the invention to provide an implantable electrical lead equipped with a tantalum electrode.

Yet another object of the invention is to provide an implantable medical device connectable to an implantable electrical lead equipped with a tantalum electrode.

These and other objects are met by the invention as defined by the accompanying patent claims.

Briefly, the present invention involves usage of tantalum or tantalum alloys for manufacturing different implantable electrodes to use in connection with medical electrical leads. The implantable electrode comprises an electrode core or substrate made of tantalum or a biocompatible tantalum alloy. At least a portion of the outer surface of the core, preferably the whole outer tissue-facing surface, is coated with a conducting layer. This conducting coating inhibits or reduces the growth and preferably also formation of a non-conducting, resistive oxide, such as tantalum pentaoxide, on the core surface. As a consequence, the coating layer combats the deterioration of the electrical and electrochemical properties of the tantalum core material that otherwise occur during use through the growth of a dielectric oxide layer on the tantalum material.

The coating material is selected from conducting nitrides, such as titanium nitride, conducting oxides, such as manganese oxide, iridium oxide or ruthenium oxide, conducting carbides, such as titanium carbide, hafnium carbide or tungsten carbide, conducting carbonitrides, such as titanium carbonitride, or conducting bromides, such as tantalum bromide or hafnium bromide. The applied conductive layer preferably presents a rough surface structure having a high effective surface area, which results in advantageous electrical properties to the implantable electrode.

The surface layer of the present invention is preferably applied to the tantalum core through physical vapor deposition or sputtering deposition.

The electrode of the present invention can be an active fixation conducting electrode, such as helix electrode, a tip electrode of a passive fixation lead, a ring electrode of an active or passive fixation lead or a defibrillation or cardioversion electrode. The electrode can be arranged on an implantable medical lead that is connectable to an implantable medical device, such as pacemaker, defibrillator or cardioverter.

The invention offers the following advantages: - A major cost reduction as compared to prior art platinum-based

• electrode materials;

Provides radiopacity by being radiopaque for X-rays, which enables visualization during implantation;

Fully biocompatible; and - Has stable electrical and electrochemical properties during use and following contact with strong oxidizing substances.

Other advantages offered by the present invention will be appreciated upon reading of the below description of the embodiments of the invention.

SHORT DESCRIPTION OF THE DRAWINGS

The invention together with further objects and advantages thereof, may best be understood by making reference to the following description taken together with the accompanying drawings, in which:

Fig. 1 is a schematic overview of a helix electrode according to the present invention; Fig. 2 is a cross-sectional view of the helix electrode of Fig. 1;

Fig. 3 is an illustration of a medical electrical lead according to the present invention;

Fig. 4 is a greatly enlarged side cross-sectional view of a distal segment of an active fixation lead according to the present invention;

Fig. 5 is a greatly enlarged side cross-sectional view of a distal segment of a passive fixation lead according to the present invention;

Fig. 6 is an illustration of an implantable medical device connectable to a medical electrical lead according to the present invention;

Fig. 7 is a SEM topography image of the rough TiN coated on Ta, showing a

■ fractal structure at a magnification of 1000Ox;

Fig. 8 is a diagram illustrating cyclic voltammetry curves of uncoated Ta exposed to PBS with an enlarged insert of the cathodic region;

Fig. 9 is a diagram illustrating cyclic voltammetry curves of uncoated Ta exposed to PBS with addition of 100 mM H2O2 with an enlarged insert of the anodic region;

Fig. 10 is a magnification of the cathodic region of the cyclic voltammetry curves of uncoated Ta exposed to PBS with addition of 100 mM H2O2 and shown in Fig. 9;

Fig. 1 1 illustrates Bode plots of uncoated Ta exposed to PBS before and after the cyclic voltammetry cycles between - 1.5 and 1.5 V vs. Ag/ AgCl; Fig. 12 illustrates Bode plots of uncoated Ta exposed to PBS with addition of 100 mM H2O2 before and after the CV cycles between -1.5 V and 1.5 V vs. Ag/ AgCl;

Fig. 13 is a diagram illustrating cyclic voltammetry curves of the rough TiN coated Ta exposed to PBS;

Fig. 14 is a diagram illustrating cyclic voltammetry curves of the rough TiN coated Ta exposed to PBS with addition of 100 mM H2O2;

Fig. 15 illustrates Bode plots of the rough TiN coated Ta exposed to PBS before and after the CV cycles between -1.5 and 1.5 V vs. Ag/ AgCl;

Fig. 16 illustrates Bode plots of the rough TiN coated Ta exposed to PBS before and after the CV cycles between - 1.5 and 1.5 V vs. Ag/ AgCl;

Fig. 17 illustrates Bode plots of uncoated tantalum exposed to PBS before and after 120 hours of simulated pacemaker pulsing; and

Fig. 18 illustrates Bode plots of the rough TiN coated Ta exposed to PBS before and after 120 hours of simulated pacemaker pulsing.

DETAILED DESCRIPTION

Throughout the drawings, the same reference characters will be used for corresponding or similar elements.

The present invention generally relates to a new type of conducting electrodes that are particularly adapted for usage in connection with implantable electrical leads. The electrode of the present invention comprises an electrode core of tantalum or a biocompatible tantalum alloy. At least a portion of the outer surface of the electrode core is then coated with a conducting surface layer having the capability of inhibiting the growth of a resistive oxide on the surface portion. Tantalum has been in clinical use since 1940 and has found a wide range of diagnostic and implantation applications. Tantalum and tantalum alloys have several physical properties that make them advantageous to use as electrode material for medical leads Tantalum has shown good biocompatibility both in hard and soft tissue, and also provides excellent mechanical properties and chemical resistance. Furthermore, tantalum is dense, ductile, very hard but easily fabricated. The transition metal is also highly conductive of electricity and has excellent resistance to corrosion.

Due to its high atomic mass and density (16.654 g/cm³), tantalum and its biocompatible alloys exhibit excellent X-ray visibility. This is an important feature as it allows a clinician or physician to visibly follow the movements and the placement of the electrode according to the present invention during implantation in a recipient patient. As a consequence, the tantalum electrode can also be utilized as a visible guiding marker at the implantation procedure, relaxing the need of using dedicated lead markers and thereby reducing the number of components of the lead.

A further advantage of tantalum and its alloys as compared to the prime lead materials today, platinum and platinum/ iridium, is that tantalum is much less expensive, approximately 20 to 30 times less expensive. This means that the total cost of manufacturing the electrical lead will be reduced markedly.

The core material can be made of a biocompatible, electrically conducting alloy material comprising tantalum, preferably comprising tantalum as a main constituent. Biocompatible relates in the present invention to a material that is non-toxic and does not cause any significant injurious effects on the subject body when being implanted. Thus, biocompatibility of the tantalum alloy implies that the electrode is able to perform its intended function, i.e. sense electrical signals from surrounding tissue and/or apply electrical signals to the surrounding tissue, with the desired degree of incorporation in the host body, without eliciting any undesirable local or systemic effects in that host.

Non-limiting examples of such tantalum alloys that can be used according to the present invention as electrode core material includes tantalum-titanium, such as Ta:Ti 20 %:70 %, and tantalum-tungsten, such as Ta:W 90 %: 10 %. It is anticipated by the present invention that the electrode core can be made of an alloy of more than one alloying material in addition to tantalum.

However, when using pure tantalum electrodes without any protective surface coating of the invention, a resistive, non-conducting oxide layer, probably tantalum pentaoxide (Ta2θs), quickly builds up on the electrode surface. Furthermore, the thickness of the dielectric oxide layer grows during use, for instance when applying pacing pulses with the electrode. Such a non-conducting oxide layer is highly undesirable for most practical applications when using the electrode in electrical medical leads connected

' to implantable medical devices (IMDs), such as pacemakers, defibrillators and cardioverters. In such a case, the non-conducting layer will interfere with the electrical signal application, thereby increasing the risk of loss of capture when providing a pacing pulse. This in turn would imply that ever higher pacing voltages are required to combat the increase in oxide layer thickness and resistivity and achieve capture. However, this drains the battery of the IMD, reducing the operation time of the lead significantly. In summary, uncoated tantalum is not suitable as electrode material together with IMDs.

As is further discussed in the experiment section, uncoated tantalum is not only unsuitable as electrode material due to the oxide formation and growth, but its electrochemical behavior can be negatively influenced by different oxidizing substances, in particular hydrogen peroxide (H2O2), released in connection with an inflammatory reaction in the host body. These and other disadvantages with the pure tantalum electrodes and Ta/Ta2θδ electrodes are solved by the present invention by providing a protective surface coating on the electrode core. In clear contrast to Ta2θs, the coating layer of the invention is a conducting layer that inhibits or at least reduces the growth and preferably formation of resistive, nonconducting oxide, in particular Ta2θs, on the outer surface of the electrode core. The surface coating of the invention works like a barrier and prevents the growth of the highly resistive oxide Ta2θs on the tantalum electrode surface.

The coating material can be provided on a portion of the outer surface of the electrode core. Thus, depending on the particular electrode design utilized, it could be possible that only a part of the outer electrode surface is exposed, during use, to the host environment, while a remaining electrode surface part is protected inside the medical lead, catheter or device housing the electrode. In such a case, only the exposed surface part needs to be coated with the conducting, oxide-inhibiting layer. However, in a preferred embodiment of the present invention, the whole outer core surface or at least a major portion thereof is provided with the protective coating of the present invention.

The conducting, oxide -inhibiting layer of the implantable electrode of the invention is made of a biocompatible and conducting oxide, nitride, carbide, bromide or carbonitride. The conducting coating layer of the invention is furthermore preferably made of a metal oxide, a metal nitride, a metal carbide, a metal bromide or a metal carbonitride. A preferred nitride coating material that can be used according to the present invention includes titanium nitride (TiN). Examples of preferred conducting oxide coatings include iridium oxide, such as iridium (IV) oxide (Irθ2), ruthenium oxide (RUO2) and manganese oxide (Mnθ2). Titanium carbide (TiC), hafnium carbide (HfC) and tungsten carbide (WC and W2C) are illustrative examples of carbide coatings that can be used according to the present invention. A metal carbontride material of the invention has the general formula of MeC_xN_1-X, where 0<x< l (x=l corresponds to a metal carbide, while a metal nitride has x=0). An example of metal (Me) that can be used is titanium, such as TiC0.5N0.5- Also a conducting bromide, such as tantalum (TaB2) or Hafnium (HfB2), can be used as coating material. The coating material can indeed be a mixture or alloy of multiple different nitrides, oxides, carbides, carbonitrides, or mixtures thereof. The outer conducting layer can be applied according to conventional techniques well-known in the art, including sputtering, physical vapor deposition, arc ion plating, etc.

The protective, conducting coating on the core electrode surface is preferably rough to present a large effective surface area. In such a case, the resulting effective surface area of the coated electrode will be many times larger than the geometric area. In a preferred embodiment, the coated conducting layer has an effective surface area resulting in surface enlargement of at least about 100 times, preferably at least about 200 times and more preferably around or at least about 300 times as compared to the original uncoated outer surface of the electrode core. The surface enlargement obtainable due to the surface roughness of the protective conducting layer provides several advantageous effects to the resulting electrode. Firstly, a large effective surface area of the electrode gives a higher capacitance for charge transfer.

Furthermore, the surface enlargement also lowers polarization after applied electric pulses, which is important when sensing electric signal from surrounding tissue with the electrode.

The surface coating of the present invention preferably has a thickness of about 2 μm to about 15 μm. Preferably the thickness is from about 3 μm to about 13 μm and more preferably in the range of from about 5 μm to about 8 μm.

Depending on the manufacturing conditions when producing the electrode of the present invention, a very thin oxide layer may spontaneously form on the tantalum core surface before applying this conducting, protective surface coating. This occurs if the tantalum core is present in an oxidizing atmosphere, such as air. The formation of such a thin oxide layer can be prevented by manufacturing the electrode in a reducing or inert environment. Alternatively, the oxide layer can be removed prior or during the coating of the tantalum core. However, the key feature of the present invention is that the conducting coating inhibits any (further) growth of a non-conducting oxide layer on the tantalum surface.

The conducting, protective surface layer of the present invention is preferably applied to tantalum core through a physical vapor deposition (PVD) method or a sputtering deposition method. The above-presented preferred coating materials of the invention can all be applied by PVD or sputtering. Furthermore, the preferred deposition procedure can be used to achieve the desired surface increase and roughness of the coating layer.

PVD and sputtering methods are preferred over chemical processes, such as chemical vapor deposition, chemical decomposition and electroplating. These methods, in clear contrast to PVD and sputtering, often requires usage of non-healthy chemicals, requiring extensive cleaning operations if the final electrodes are to be implanted in an animal or human body.

The present invention can be applied to different types of electrodes traditionally employed for electrical leads. For example, the electrode can be an active electrode tip, such as an active fixation electrode. Helices and hooks are non-limiting examples of such structures that have the dual function of both actively fixing the lead into a body tissue, such as endocardium or myocardium, and operate as a conducting electrode. Another electrode example is the electrodes of a passive electrode tip in a lead. In this case, another structure of the lead tip, such as a collar, a tine assembly or a fin assembly, passively fixes the lead to the body tissue. The electrode does then not actively participate in this tissue fixation. Bipolar and multipolar leads are also employed in the art. In such a lead, the tip electrode as described above is complemented with one or more separate ring electrodes, typically positioned at a distance from the lead tip. The electrode can also be a defibrillation or cardioversion electrode for applying a (high voltage) defibrillation or cardioversion shock.

Fig. 1 illustrates an active fixation electrode 28 of the present invention in the form of a helix. The portion of the helix 28 denoted by A, i.e. the end portion of the helix 28 that is to be inserted into the myocardium can be coated, while the remaining portion is left uncoated or coated differently. In this context, the partly coating could be TiN (or a similar material, see above), while the remaining electrode surface is uncoated.

Fig. 2 is a cross-sectional view of the tantalum wire employed for forming the active, helical fixation electrode of Fig. 1. As is seen in the figure, the implantable electrode comprises an inner electrode core 60 of tantalum or the biocompatible tantalum alloy. The outer surface 65 of the core is coated with the conducting layer 70 inhibiting the growth of the non-conducing oxide on the core surface 65.

The present invention will now be further illustrated in connection with implementation in different electrical leads. For the purposes of this specification, the term "electrical lead" is used herein in its broadest sense and includes a stimulation lead, a sensing lead, a combination thereof or any other elongated member, such as a catheter, which may be introduced into a body and has at least one electrode according to the present invention. Examples of such leads include endocardial and epicardial leads.

Fig. 3 schematically illustrates a lead 10 according to the present invention. The lead 10 has a flexible, elongated lead body 12 covered by an insulative sleeve made of, for example, polyure thane, silicone rubber or a mixture thereof. Terminal assembly 14 is provided at the proximal end for coupling the lead 10 to an implantable medical device, IMD, (not shown.) The terminal assembly 14 has sealing rings 16 and terminal pin 18, all of a type known in the art. An optional anchoring sleeve 20 (shown partially in cross-section) may also be provided for suturing the lead body 12 to body tissue following implantation. The anchoring sleeve 20 and the terminal assembly 14 are preferably fabricated from silicone rubber, although they may also be constructed of any other suitable biocompatible material known in the art.

The lead 10 may also be connected to a stylet guide 22 through the terminal pin 18. This stylet guide 22 is used together with a stylet assembly 24 for imparting stiffness to the lead 10 during placement. The stylet guide 22 and the stylet assembly 24 are typically discarded after use and before connection of the terminal pin 18 to the IMD.

An electrode and fixation assembly 26 is provided at the distal end of the lead body 12. The electrode and fixation assembly 26 is, in the disclosed embodiment, of the bipolar and active type and has a helix 28 at its distal end and a ring electrode 30 spaced proximally back from the distal end. As will be appreciated by those of ordinary skill in the art, the helix 28 and the ring electrode 30 are coupled to separate, insulated lead conductors which extend along the length of the lead body 12. Lead conductors are preferably configured as concentric multi-filar coils of MP35N, 35N LT or any other suitable alloy, such as a platinum-iridium alloy. This configuration allows for a longitudinal lumen to exist along the length of the lead body 12, such that the stylet may be received therein.

Either of the helix electrode 28 and the ring electrode 30, and more preferably both electrodes 28, 30, are according to the present invention, i.e. surface coated tantalum or tantalum alloys.

Fig. 4 illustrates a greatly enlarged cross-sectional side view of the distal portion of the lead body 12 and the electrode and fixation assembly 26 of Fig. 3. As seen, the lead body 12 has an outer flexible insulative sheath 32 made of, for instance silicone rubber, polyurethane, a mixture thereof, or the like. The outer insulative sheath 32 covers a first coiled conductor 34. The conductor 34 extends along through the lead body 12 and terminates at its distal end where it is electrically coupled, for example by spot or laser welding, to a crimp sleeve 36. The crimp sleeve 36, in turn, is in electrical connection with a ring electrode 30, which is made according to the present invention in tantalum or tantalum alloy with a conducting oxide-inhibiting surface coating.

Partially engaged between the ring electrode 30 and the helix 28 is a ring/ spacer assembly 31 that is coupled to a tip/ring spacer 40, which is typically made of silicone rubber. In addition to establishing a predetermined distance between the ring electrode 30 and the helix 28, the tip/ ring spacer

40 functions to define a substantially cylindrical chamber in which the remaining components are disposed as well as to define the outer surface of the electrode and fixation assembly 26. In the disclosed embodiment, the tip/ ring spacer 40 has dimensions such that a constant lead body diameter is maintained between the helix 28 and the ring electrode 30.

Extending along the length of the lead body 12 through the crimp 36, the ring electrode 30, the ring/ spacer assembly 31 and the tip/ring spacer 40 is a second coiled conductor 42, which is insulated from the outer coiled conductor 34 by an inner insulative sheath 44 which, like the outer sheath 32 can be made of silicone rubber, polyure thane, a mixture thereof, or the like. The inner conductor 42 terminates at a substantially cylindrical crimp bus 46. The crimp bus 46, in turn is coupled to the helix 28. Located distal to the crimp bus 46 is an optional indicator ring 47 to provide a radiopaque indication of how far extended the helix 28 is from the lead body 12. The helix 28 has the dual function of fixing the lead to the myocardium and functions as a pacing/ shocking electrode. The helix 28 is made according to the present invention in tantalum or tantalum alloy with a conducting surface coating. In an alternative embodiment of an active fixation lead according to the present invention, the tantalum ring electrode is omitted. In other words, the lead is of the unipolar type. The tantalum electrode is then the active helix fixation electrode or another type of active tantalum fixation electrode.

The implantable electrode of the invention can also be used in other active fixation leads than the one illustrated in Fig. 4, i.e. having another set of distal lead components.

Fig. 5 illustrates an enlarged cross-sectional view of an electrical lead of passive fixation type equipped with tantalum electrodes 30, 38 according to the present invention. The lead body 12 has an outer flexible insulative sheath 32 covering a first coiled conductor 34 extending along the lead body 12 and terminating at a tantalum ring electrode 30 or electrically coupled to the ring electrode through a crimp sleeve (not illustrated) .

A second inner coiled conductor 42 is electrically insulated from the first conductor 34 by an inner insulative sheath 44. The inner conductor 42 is electrically connected to a tantalum tip electrode 38 according to the present invention. In Fig. 5, the electrode coating 70 in the form of the conducting oxide-inhibiting layer, such as rough TiN, is schematically illustrated. The lead tip may also optionally be equipped with a steroid plug 54, the use of which is well known in the art. Passive fixation of the lead at a correct position in a patient body is achievable by a tine assembly 50.

It is anticipated by the present invention that the lead equipped with electrodes 30, 38 according to the present invention and partly illustrated in Fig. 5 may alternatively by of a unipolar type. In such a case, the ring electrode 30 and its associated coiled conductor 34 can be omitted.

The implantable electrode of the invention can also be used in other passive fixation leads than the one illustrated in Fig. 5, i.e. having another set of distal lead components. The lead of the invention described and disclosed above can, in operation, be connected to an IMD. The IMD could be a pacemaker, cardiac defibrillator or cardioverter. In such a case the IMD has at least one connected lead with electrodes and/or sensors for performing the therapy and/or monitor function to the heart. However, the present invention is not limited to cardiac-associated IMDs and leads but may also be practiced with other electrical lead implantable for performing therapy and/ or diagnosing in other body positions besides the heart, such as a neurological or muscle stimulating electrical lead.

Fig. 6 schematically illustrates an IMD 100, exemplified as a pacemaker, connectable with a medical lead 10 having at least one implantable electrode according to the present invention.

EXPERIMENTS

Electrochemical behavior of Ta and rough TiN coated Ta exposed to phosphatebuffered saline solution (PBS) with and without addition of H2O2 were investigated, using cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) and simulated pacemaker pulses.

Material, surface preparation

The Ta foil of 1.0 mm thickness and 99.95 % purity was purchased from Alfa

Aesar, Johnson Matthey. The foil was cut into smaller pieces, 10 mm x 10 mm. Some of the samples were sent to Heraeus (Hanau, Germany) for coating of rough TiN by physical vapor deposition. The Ta samples with the rough TiN coating were investigated in as-received condition. Fig. 7 shows a SEM micrograph of the rough surface topography of the TiN coating 70. A large effective surface area is preferred for pacemaker electrodes since that gives a higher capacitance for charge transfer and also a lower polarization, which is important for sensing of the cardiac signals [I]. Based on the measured capacitive values the surface enlargement with the rough TiN is about 300 times compared with smooth TiN. The thickness of the rough TiN coating is 6 μm. The Ta samples without coating were polished with SiC paper down to grit 1200 and rinsed with deionized water and degreased with isopropanol. A geometric surface area of 0.5 cm² was exposed to the solution during the electrochemical evaluation. 5

Electrolyte and electrochemical cell

A standard PBS solution purchased from VWR International was used for the electrochemical evaluation. The solution composition was 8000 mg/L NaCl, 200 mg/L KCl, 1150 mg/L Na₂HPO₄ and 200 mg/L KH₂PO₄, with

LO similar ion concentrations and pH as blood. Furthermore, addition of 100 mM H₂O₂ to the PBS electrolyte was employed in selected experiments to simulate the conditions of an inflammatory reaction where H₂O₂ is produced. Such a solution has been used previously to investigate corrosion resistance of implant materials [2, 3].

L5

, The electrochemical cell used was a standard three-electrode electrochemical cell, with a saturated Ag/ AgCl reference electrode, a Pt mesh as the counter electrode, and the sample as the working electrode. The electrolyte volume was 300 ml. All the electrochemical measurements were performed at room

20 temperature.

CV and EIS measurements

A versatile multichannel potentiostat (VMP2) was used for the CV and EIS measurements, and the instrument was controlled by a computer with EC- .5 lab^® v9.30 software. Both the instrument and the software were supplied by

Bio-Logic SAS. The CV measurements were performed to investigate electrochemical reactions that may occur on the surface of the material, and the influence of formation of insulating oxides on the surface. The CV was performed between -1.5 to 1.5 V vs. Ag/ AgCl reference electrode for 10 or 15

30 cycles. This potential window was chosen to cover eventual oxygen evolution and hydrogen evolution reactions. This is also within the potential range relevant for pacemaker pulses [4]. The potential sweep rate used for the cycling was 100 mV/s, which is relatively fast, but much slower than the transient process in the pacemaker pulses. The EIS measurements were performed before and after the CV cycles to characterize the interfacial electrochemical behavior, and to detect the changes due to the alternating polarization during the CV cycles. The EIS measurements were performed at open circuit potential conditions, over a frequency range from 10 kHz down to 1 mHz. The ac perturbation amplitude was 15 mV.

Being an accelerated test, the potential range and the sweep rate of the CV measurements are not representative for the pacing conditions. During normal pacemaker pulses, the potential may be much lower than - 1.5 V, but the time for the transient process is much shorter than that in the CV measurements. Therefore, additional simulated pacing pulses were applied to the material, and the EIS measurements were performed before and after the pulses to evaluate the eventual change caused by the simulated pulses (below).

Cyclic voltammetry for uncoated Ta

The CV curves recorded for the uncoated Ta sample exposed to PBS are shown in Fig. 8, the curve for the very first cycle was not included because of large current peaks apparently due to oxidation and reduction of some absorbed surface species. It can be seen that enhanced current occurs both at anodic potentials above ca. 0.9 V, and at cathodic potentials below approximately - 0.7 V, and the current increases drastically towards 1.5 V or -1.5 V (limits of the CV performed). The results indicate that some oxidation reaction occurs at the high anodic potentials, probably the oxygen evolution, and some reduction reaction occurs at the low cathodic potentials, most likely the hydrogen evolution, occurring on the Ta surface.

Moreover, the current density in both the anodic and cathodic potential regions decreases with increasing number of the CV cycles. This indicates the formation of a highly resistive anodic oxide film on the surface. The oxide film formed on the surface is most likely Ta2θs. It is well known that Ta forms an anodic oxide layer on the surface and the oxide layer thickness is directly proportional to the anodizing voltage, about 10 - 20 A per volt [5-7]. In the cathodic region, the decrease in current density with the number of cycles suggests that the cathodic reaction rate decreases due to the increasing resistance of the oxide film growing on Ta surface. This also implies that, under current experimental conditions, uncoated Ta in PBS solution is not subjected to a great extent of hydrogen uptake, because of the formation of a surface oxide layer on Ta.

Fig. 9 shows the CV curves of Ta exposed to PBS with addition of 100 mM H2O2, and the anodic region is magnified for clarity. The first cycle exhibits the highest current density, and a shape different from the following cycles.

The magnification of the cathodic region of the CV curves of Ta exposed to

PBS with addition of 100 mM H2O2 is displayed in Fig. 10, showing detailed variation in the first several cycles (the arrows indicate increasing number of cycles). The first cycle (C l) shows the highest current density in the cathodic region. The current density decreases drastically in cycle 2 (C2), but then increases until it reaches a stable stage, see cycles 5- 10.

In general, these CV curves are quite different from those obtained in PBS without H2O2 (Fig. 8). With the H2O2 addition, the enhanced current occurs at relatively small anodic and cathodic potentials; two anodic current peaks appear in the anodic region below 1.5 V (Fig. 9) and one cathodic current peak appears around - 1.0 V, especially in the first several cycles. Moreover, there is a large current loop in the anodic region (Fig. 9) even after several cycles. The current peaks are probably associated with oxidation and reduction reactions involving H2O2 on the Ta surface. Metal surfaces are known to catalyze decomposition of H2O2. Since the H2O2 was only added prior to the electrochemical measurements, its concentration will decrease with the consumption of the H2O2, and hence the reaction rate, especially in the cathodic region. The decomposition of H2O2 generates O2 on the Ta surface, which will facilitate the oxidation of Ta and oxide formation in the anodic region. This may explain the decreasing current density with increasing number of CV cycles as discussed above. The results obtained show clearly that the addition of H2O2 influences the electrochemical behavior of uncoated Ta until all hydrogen peroxide is consumed.

Electrochemical impedance spectroscopy for uncoated Ta The EIS spectra of uncoated Ta before and after ten CV cycles are shown in

Bode plots in Fig. 1 1. It is clear that the impedance in the low frequency region is greatly increased after the CV cycles. Despite of some distortion in the phase angle data, the spectra can be fitted to the simplest equivalent circuit describing the electrode electrolyte interface, i.e. parallel interfacial capacitance and polarization resistance, in series with the electrolyte resistance. As usual, for the spectra fitting, a constant phase element is used instead of the capacitance to account for non-ideal capacitive behavior of the interface. The numerical values of the electrical components, the polarization resistance R_p and interfacial capacitance (constant phase element, CPE), obtained from the spectra fitting are summarized in Table 1 below, where all data are mean values of at least triplicate samples. The polarization

• resistance increased by two orders of magnitudes after the CV cycles. The greatly increased polarization resistance and decreased capacitance are most likely due to the formation of a highly resistive oxide, in agreement with the CV measurements. It was reported that Ta could develop an extremely high- impedance surface layer, which can insulate the electrode to the point of nonfunction [8].

Fig. 12 shows the EIS spectra of Ta exposed to PBS with addition of 100 mM H2O2. In this case, the polarization resistance increased by one order of magnitude after the CV cycles. Again, the EIS results are in agreement with the CV measurements, suggesting the formation of highly insulating oxide film on the uncoated Ta surface during the CV cycles. On the other hand, the polarization resistance of Ta exposed to PBS with the addition of 100 mM H2O2 after the CV cycles is slightly lower as compared to the case without

H2O2 addition. It was observed previously that H2O2 may cause an enhanced corrosion attack on the metal surface [9], and that may be the reason for a lower polarization resistance in the case with the H2O2 addition. In short, the results indicate that Ta forms a highly resistive anodic oxide film on the surface during the alternating anodic and cathodic polarization. Moreover, the addition of H2O2 influences the electrochemical behavior and 5 the oxide formation on the surface of Ta, leading to a slightly decreased corrosion resistance.

Cyclic voltammetry for Ta surface-coated with rough TiN

Fig. 13 shows the CV curves of the rough TiN coated Ta exposed to PBS. The

L 0 current density level is about two order of magnitudes higher compared to Ta without coating (Fig. 8), due to a much larger surface area of the rough TiN coating. However, the shape of the CV curves is similar to that of the uncoated Ta, and the current density also decreases with increasing number of the CV cycles (indicated by arrows). The results suggest that, under the

L 5 CV conditions, the TiN coating is subjected to some oxidation process at anodic potentials, leading to the formation of a resistive oxide film on the surface. It was previously reported that oxygen atoms penetrate into the TiN coating and replace nitrogen atoms to form a Tiθ2-like oxide [10], most likely according to [H]:

20

2TiN + 2O₂ → 2TiO₂ +N₂ (1)

In the cathodic region of the CV cycles, the current density decreases with increasing number of the CV cycles. The increased resistance of the oxide

25 film formed in the anodic region could be one reason for this observation.

Moreover, the conductivity of TiN may also decrease at cathodic potentials due to the formation of TiH_x, which was reported to exhibit a lower conductivity than TiN [12]. However, it seems that the oxide film formed on the surface in the anodic region is not reduced in the cathodic region. Even

30 the formation of a hydrated oxide film would give a higher conductivity [12-

14], and hence an increased current. Fig. 14 shows the CV curves of the rough TiN coated Ta exposed to PBS with addition of 100 mM H2O2. The oxidation at high anodic potentials and reduction at low cathodic potentials are evident, but the current density decreases with increasing number of cycles, which indicates the formation of a resistive oxide film on the surface. In general, the CV curves are different from those of uncoated Ta exposed to PBS with addition of 100 mM H2O2 (Fig. 9), but similar to those for the rough TiN coated Ta without H2O2 (Fig. 13). With the H2O2 addition, the current density in the first cycle is considerably higher than in the following cycles, but the current peaks are not so pronounced as in the case for uncoated Ta. The results indicate that the addition of H2O2 also has much less influence on the electrochemical behavior of the rough TiN coated Ta compared to the uncoated Ta.

Electrochemical impedance spectroscopy for Ta surface-coated with rough TiN The EIS spectra for the TiN coated Ta exposed to PBS are shown in Fig. 15.

These CV cycles have not caused any significant change, and based on the same spectra fitting as above, the polarization resistance remains to be about the same level after the CV cycles, see Table 1 below which lists electrical components from spectra fitting for Ta and the TiN/Ta exposed to PBS solution with and without H2O2, before and after the CV cycles.

Table 1 - Electrical components

Material/ Electrolyte Rp (Ωcm²) Q (F/cm2) α

Ta/ PBS Before CV 8. IxIO⁴ 4.5xlO-⁵ 0.92

After CV 6.2xlO⁶ 1.6x10-5 0.92

Ta/PBS+H₂O₂ Before CV 1.4xlO⁵ 3. IxIO-⁵ 0.95

After CV 7.3x10⁵ 1.5x10-5 0.94

TiN-coated Ta/ PBS Before CV 1.2xlO⁴ 0.022 0.89

After CV 1.0x10⁴ 0.015 0.88

TiN-coated Ta/PBS+H₂O₂ Before CV 2.6x10³ 0.019 0.85

After CV 4.6x10³ 0.013 0.84

Rp - polarization resistance Q - constant phase element α - exponential factor of the constant phase element

The results for the rough TiN coated Ta exposed to PBS with addition of 100 mM H2O2 are similar to those without H2O2 addition, indicating a small change due to the CV cycles. Fig. 16 shows the EIS spectra for the rough TiN coated Ta in PBS with addition of H2O2. Clearly the rough TiN coating on Ta drastically altered the electrochemical behavior and reduced the influence of

H2O2. The influence of surface roughness on the electrochemical behavior of pacemaker electrodes has been previously explained by the ohmic drop inside the pores and reduced potential penetration into the pores [15].

Pacemaker pu lses

To simulate pacing conditions on the electrode material, a pacemaker was connected to the three-electrode cell. The pacemaker used for the pulsing was an Identity^® ADx DR model 5380. The pacing settings were set to, base rate 60 min ¹, cathodic pulse amplitude (negative) 7,5 V and the pulse width to 0.4 ms.. Cathodic pulse amplitudes are used in all pacemakers today [16]. A pacemaker was placed in a Faraday cage, with a special costume made connector pin inserted in the pacemaker connector enabling the sample to be the working electrode and the Pt mesh to be the counter electrode.

Pacemaker pulses for uncoated Ta

Fig. 17 shows the EIS spectra for the uncoated Ta exposed to PBS before and after 120 hours of the simulated pacemaker pulsing. The results show that the polarization resistance increased by three orders of magnitudes after the pulsing. The high polarization resistance after the pulsing indicates that a highly resistive surface oxide has formed on Ta during the pulsing. In agreement with the CV results, these studies indicate that uncoated Ta is not a suitable electrode material.

Pacemaker pulses for Ta surface-coated with rough TiN

Fig. 18 shows the EIS spectra of the rough TiN coated Ta exposed to PBS before and after 120 hours of the simulated pacemaker pulsing. The polarization resistance remains basically unchanged, in some cases even slightly decreased, after the pulsing. This stability is desirable for the pacemaker electrodes, and the slight decrease in polarization resistance upon the pacemaker pulsing is probably beneficial with respect to charge transfer capacity of the electrode. In agreement with the CV measurements, the results indicate that Ta is a suitable substrate material for the rough TiN coated pacemaker electrodes.

It will be understood by a person skilled in the art that various modifications and changes may be made to the present invention without departure from the scope thereof.

REFERENCES

[1] Norlin et al., Leygraf, Biomol. Eng., 19, 67 (2002)

[2] Pan et al., J. Biomed. Mater. Res., 28, 1 13 (1994)

[3] Pan et al., J. Biomed. Mater. Res., 30, 393 (1996)

[4] Norlin et al., J. Electrochem. Soc, 152, J7 (2005)

[5] Robblee et al., J. Biomed. Res., 17, 327 (1983)

[6] Guyton and Hambrecht, Science, 181, 74 (1973)

[7] Kerrec et al., Electrochim. Acta, 40, 719 (1995)

[8] Stokes, Proc. of the IEEE, 84, 457 (1996)

[9] Pan et al., J. Biomed. Mater. Res., 28, 1 13 (1994)

[10] Azumi et al., Corros. Sά., 40, 1363 (1998) [11] Milosev et al., Thin solid films, 303, 246 (1997)

[12] Azumi et al, J. Electrochem. Soc, 149, B422 (2002)

[13] Dyer and Leach, Electrochem. Acta, 23, 1387 (1978)

[14] Ohtsuka et al., J. Electrochem. Soc, 134, 2406 (1987)

[15] Norlin et al., J. Electrochem. Soc, 152, JI lO (2005)

[16] Norlin, Investigation of electrochemical properties and performance of stimulation/ sensing electrodes for pacemaker applications, Doctoral thesis, Royal Institute of Technology, Stockholm, Sweden (2005). ISBN 91-7283-994-5

[17] Brummer and Turner, Bioelectrochem. Bioenergetics, 2, 13 (2975)

[18] Johnson et al., J. Biomed. Mater. Res., 11, 637 (1977)

Claims

1. An implantable electrode (28, 30, 38) comprising: an electrode core (60) of tantalum or a biocompatible tantalum alloy; and - a conducting layer (70) coated on at least a portion of an outer surface (65) of said electrode core (60) and inhibiting growth of a resistive oxide on said at least a portion of said outer surface (65), said conducting layer (70) is made of at least one of the group consisting of a biocompatible conducting oxide, nitride, bromide, carbide and carbonitride.

2. The implantable electrode according to claim 1, wherein said electrode core (60) is made of tantalum.

3. The implantable electrode according to claim 1, wherein said biocompatible tantalum alloy is a biocompatible alloy between tantalum and

^■ at least one of titanium and tungsten.

4. The implantable electrode according to any of the claims 1 to 3, wherein said conducting layer (70) is coated on said outer surface (65) of said electrode core (60).

5. The implantable electrode according to any of the claims 1 to 4, wherein said conducting layer (70) is made of at least one of the group consisting of a biocompatible conducting metal oxide, metal nitride, metal bromide, metal carbide and metal carbonitride.

6. The implantable electrode according to claim 5, wherein said conducting layer (70) is made of titanium nitride.

7. The implantable electrode according to claim 5, wherein said conducting layer (70) is made of a metal oxide selected from the group consisting of iridium oxide, ruthenium oxide and manganese oxide.

8. The implantable electrode according to claim 5, wherein said conducting layer (70) is made of titanium carbonitride.

9. The implantable electrode according to claim 5, wherein said conducting layer (70) is made of a metal bromide selected from the group consisting of tantalum bromide and hafnium bromide.

10. The implantable electrode according to claim 5, wherein said conducting layer (70) is made of a metal carbide selected from the group consisting of titanium carbide, hafnium carbide and tungsten carbide.

11. The implantable electrode according to any of the claims 1 to 10, wherein said conducting layer (70) has an effective surface area resulting in a surface enlargement of at least about 100 times, preferably at least about 200 times and more preferably at least about 300 times as compared to said outer

• surface (65) of said electrode core (60).

12. The implantable electrode according to any of the claims 1 to 11, wherein said conducting layer (70) inhibits growth of tantalum pentaoxide on said at least a portion of said outer surface (65).

13. The implantable electrode according to any of the claims 1 to 12, wherein said conducting layer (70) has a thickness of 2 μm to 15 μm, preferably 3 μm to 13 μm and more preferably 5 μm to 8 μm.

14. The implantable electrode according to any of the claims 1 to 13, wherein said implantable electrode (28) is an active fixation, conducting electrode (28).

15. The implantable electrode according to claim 14, wherein said active fixation, conducting electrode (28) has a general helix shape.

16. The implantable electrode according to any of the claims 1 to 13, wherein said implantable electrode (28) is a tip electrode (28) of an implantable passive fixation electrical lead (10).

17. The electrode according to any of the claims 1 to 13, wherein said implantable electrode (30) is a ring electrode (30) of an implantable electrical lead (10).

18. An implantable electrical lead (10) comprising at least one implantable electrode (28, 30, 38) as defined in any of the claims 1 to 17.

19. The lead according to claim 18, wherein said lead (10) is a cardiac lead (10).

20. An implantable medical device (100) connected to an implantable electrical lead (10) as defined in claim 18 or 19.