Dissolving barrier of the prosthesis that stimulates tissue

Abstract

Dissolution barriers for prostheses for stimulating tissue are provided herein. As further described below, the prosthesis for stimulating tissue includes a stimulation assembly including an elongated insulating carrier member and a plurality of electrode contacts disposed along the carrier member. A continuous dissolution barrier is disposed on a surface of the stimulation assembly so as to substantially encapsulate / surround the plurality of electrode contacts and the carrier member. The continuous dissolution barrier is configured to inhibit in situ dissolution of the plurality of electrode contacts.

Description

Technical Field

[0001] Certain aspects presented herein generally relate to dissolving barriers of prostheses for stimulating tissue. Background Art

[0002] There are several types of medical devices / implants that operate by delivering electrical (current) stimulation to the recipient's nerves, muscles, tissue fibers, or other cells. These medical devices, sometimes referred to herein as tissue-stimulating prostheses, typically deliver electrical stimulation to compensate for the recipient's defect. For example, when a recipient experiences sensorineural hearing loss due to loss or destruction of cochlear hair cells (which convert acoustic signals into nerve impulses), a tissue-stimulating hearing prosthesis, such as a cochlear implant, is typically recommended. An auditory brainstem stimulator is another type of tissue-stimulating hearing prosthesis that may be recommended when a recipient experiences sensorineural hearing loss due to damage to the auditory nerve. Summary of the invention

[0003] In one aspect, a prosthesis for stimulating tissue is provided. The prosthesis for stimulating tissue includes: a stimulation assembly including an elongated carrier member and a plurality of electrode contacts spaced apart along the elongated carrier member, the plurality of electrode contacts configured to perform at least one of providing a current signal or sinking a current signal; and a dissolving barrier encapsulating the stimulation assembly, wherein the continuous dissolving barrier is configured to inhibit in situ dissolution of the plurality of electrode contacts.

[0004] In another aspect, a device is provided. The device includes: an implantable carrier member formed of an insulating material and having a central longitudinal axis and an outer surface; a plurality of electrode contacts distributed along the longitudinal axis on the outer surface of the carrier member for applying electrical stimulation signals to adjacent neural tissue; and a continuous dissolving barrier laminated to each of the plurality of electrode contacts and the carrier member.

[0005] In another aspect, a method is provided. The method includes: providing a plurality of electrode contacts arranged in an elongated array; encapsulating the plurality of electrode contacts in an implantable carrier member formed of an insulating material, wherein the plurality of electrode contacts extend along a longitudinal axis of the carrier member and are positioned at an outer surface of the carrier member; exposing a surface of each of the plurality of electrode contacts for applying an electrical stimulation signal to adjacent neural tissue; and laminating a continuous dissolving barrier over each of the plurality of electrode contacts and the carrier member. BRIEF DESCRIPTION OF THE DRAWINGS

[0006] Embodiments of the present invention are described herein with reference to the accompanying drawings, in which:

[0007] Figure 1A is a schematic diagram illustrating a cochlear implant according to certain embodiments presented herein;

[0008] Figure 1B yes Figure 1A A block diagram of a cochlear implant;

[0009] Figure 2A is a schematic cross-sectional view of a dissolving barrier applied to a stimulation component according to certain embodiments presented herein.

[0010] Figure 2B It is set on the electrode contact Figure 2A A magnified view of a portion of the dissolution barrier;

[0011] Figure 3A is a schematic diagram illustrating variables for calculating different parameters of a dissolution barrier according to certain embodiments presented herein;

[0012] Figure 3B is a schematic diagram illustrating current shunting between electrode contacts for barrier materials with inappropriate conductivity.

[0013] Figure 4 is a schematic diagram illustrating a self-assembled monolayer forming a dissolution barrier according to certain embodiments presented herein;

[0014] Figure 5 is a schematic diagram of a spinal cord stimulator according to certain embodiments presented herein; and

[0015] Figure 6 is a flow chart of a method according to certain embodiments presented herein. DETAILED DESCRIPTION

[0016] Dissolution barriers for use with prostheses for stimulating tissue are provided herein. As further described below, the prosthesis for stimulating tissue includes a stimulation assembly including an elongated insulating carrier member and a plurality of electrode contacts disposed along the carrier member. A continuous dissolution barrier is disposed on a surface of the stimulation assembly so as to encapsulate / surround (e.g., substantially cover with potential small defects) the plurality of electrode contacts and the carrier member. The continuous dissolution barrier is configured to inhibit in situ dissolution of the plurality of electrode contacts.

[0017] As described above, there are many types of prostheses that stimulate tissue, which deliver stimulation signals (current signals) to compensate for the recipient's defects. For ease of explanation only, the embodiments presented herein are described herein mainly with reference to one type of prosthesis that stimulates tissue, i.e., a cochlear implant. However, it should be understood that the technology presented herein can be used with other prostheses that stimulate tissue, including, for example, auditory brainstem stimulators, implantable pacemakers, spinal cord stimulators, deep brain stimulators, motor cortex stimulators, sacral nerve stimulators, pudendal nerve stimulators, vagus nerve / vagus nerve stimulators, trigeminal nerve stimulators, retinal or other visual prostheses / stimulators, occipital cortex implants, diaphragm (diaphragm) pacemakers, analgesic stimulators, other nerves or neuromuscular stimulators, etc.

[0018] Figure 1A is a schematic diagram of an exemplary cochlear implant 100 according to aspects presented herein, and Figure 1B is a block diagram of a cochlear implant 100. For ease of illustration, Figure 1A and 1B .

[0019] The cochlear implant 100 includes an external component 102 and an internal / implantable component 104. The external component 102 is attached directly or indirectly to the body of the recipient and typically includes an external coil 106 and a magnet (not shown in FIG. 1 ) that is typically fixed relative to the external coil 106. The external component 102 also includes one or more input elements / devices 113 for receiving input signals at a sound processing unit 112. In this example, the one or more input devices 113 include a sound input device 108 configured to capture / receive input signals (e.g., a microphone positioned by the recipient's pinna 110, a telecoil, etc.), one or more auxiliary input devices 109 (e.g., an audio port such as a direct audio input (DAI), a data port such as a universal serial bus (USB) port, a cable port, etc.), and a wireless transmitter / receiver (transceiver) 111, each of which is positioned in, on, or near the sound processing unit 112.

[0020] The sound processing unit 112 also includes, for example, at least one battery 107, a radio frequency (RF) transceiver 121, and a processing module 125. The processing module 125 may include a plurality of elements, including a sound processor 131.

[0021] exist Figure 1A and Figure 1BIn the example of the embodiment of the present invention, the sound processing unit 112 is a behind-the-ear (BTE) sound processing unit that is configured to be attached to and worn adjacent to the recipient's ear. However, it should be understood that embodiments of the present invention may be implemented with sound processing units having other arrangements, such as with an outside-the-ear (OTE) sound processing unit (i.e., a component having a generally cylindrical shape and configured to be magnetically coupled to the recipient's head), a mini or micro BTE unit, an in-the-ear unit configured to be located in the recipient's ear canal, a body-worn sound processing unit, etc.

[0022] return Figure 1A and Figure 1B The implantable component 104 includes an implant (main module) 114, a lead region 116, and an intracochlear stimulation assembly 118, all of which are configured to be implanted beneath the skin / tissue (tissue) 105 of a recipient. The implant 114 typically includes a sealed housing 115 in which an RF interface circuit 124 and a stimulator unit 120 are disposed. The stimulator unit 120 includes, among other things, one or more current sources on an integrated circuit (IC).

[0023] The implant body 114 also includes an internal / implantable coil 122, which is typically external to the housing 115 but is accessible via a sealed feedthrough ( Figure 1B ) is connected to the RF interface circuit 124.

[0024] As described above, the cochlear implant 100 includes an external coil 106 and an implantable coil 122. The coils 106 and 122 are typically antenna coils of wire, each consisting of multiple turns of electrically insulated single or multiple strands of platinum or gold wire. Typically, a magnet is fixed relative to each of the external coil 106 and the implantable coil 122. The magnets fixed relative to the external coil 106 and the implantable coil 122 facilitate operational alignment of the external coil with the implantable coil. This operational alignment of the coils 106 and 122 enables the external component 102 to transmit data, and possibly power, to the implantable component 104 via a tightly coupled wireless link formed between the external coil 106 and the implantable coil 122. In some examples, the tightly coupled wireless link is a radio frequency (RF) link. However, various other types of energy transmission, such as infrared (IR), electromagnetic, capacitive, and inductive transmission, can be used to transmit power and / or data from the external component to the implantable component, thus, Figure 1B Only one example arrangement is shown.

[0025] As described above, the sound processing unit 112 includes a processing module 125. The processing module 125 is configured to convert an input audio signal into a stimulation control signal 136 for stimulating a first ear of a recipient (i.e., the processing module 125 is configured to perform sound processing on the input audio signal received at the sound processing unit 112). In other words, the sound processor 131 (e.g., one or more processing elements implementing firmware, software, etc.) is configured to convert a captured input audio signal into a stimulation control signal 136, which represents a stimulation signal for delivery to the recipient. The input audio signal that is processed and converted into a stimulation control signal may be an audio signal received via the sound input device 108, a signal received via the auxiliary input device 109, and / or a signal received via the wireless transceiver 111.

[0026] exist Figure 1B In an embodiment of the present invention, the stimulation control signal 136 is provided to the RF transceiver 121, which transmits the stimulation control signal 136 (e.g., in an encoded manner) transcutaneously to the implantable component 104 via the external coil 106 and the implantable coil 122. That is, the stimulation control signal 136 is received at the RF interface circuit 124 via the implantable coil 122 and provided to the stimulator unit 120. The stimulator unit 120 is configured to utilize the stimulation control signal 136 to generate a stimulation signal (e.g., a current signal) for delivery to the recipient's cochlea via the stimulation component 118. In this manner, the cochlear implant 100 electrically stimulates the recipient's auditory nerve cells in a manner that causes the recipient to perceive one or more components of the input audio signal, bypassing the missing or defective hair cells that normally convert acoustic vibrations into neural activity.

[0027] More specifically, as described above, stimulation assembly 118 is configured to be at least partially implanted in a recipient's cochlea 137. Stimulation assembly 118 includes a plurality of longitudinally spaced intracochlear electrical contacts (electrode contacts or electrodes) 126 that collectively form an electrode contact array 128 that is configured to deliver electrical stimulation signals (current signals) generated based on stimulation control signals 136 to the recipient's cochlea, for example. In some examples, electrode contacts 126 may also be used to receive stimulation signals from the recipient's cochlea.

[0028] Figure 1A One particular arrangement is shown in which the stimulation assembly 118 includes twenty-two (22) intracochlear electrode contacts 126, which are labeled electrode contacts 126(1) through 126(22). It should be understood that the embodiments presented herein may be implemented in alternative arrangements having different numbers of intracochlear electrode contacts.

[0029] As shown, the intracochlear electrode contacts 126(1)-126(22) are disposed in an elongated carrier member 134. The carrier member 134 has a central longitudinal axis and an outer surface. The carrier member 134 is formed of a non-conductive (insulating) material, such as silicone or other elastomeric polymers. Thus, the carrier member 134 electrically isolates the intracochlear electrode contacts 126(1)-126(22) from each other. Figure 1B As shown, the intracochlear electrode contacts 126 ( 1 )- 126 ( 22 ) are each separated from one another by segments / segments of the carrier member 134 .

[0030] Stimulation assembly 118 extends through an opening in the recipient's cochlea (eg, cochleostomy, round window, etc.) and has a proximal end that is coupled to lead region 116 and ( Figure 1B The carrier member 134 and the lead region 116 each include a plurality of conductors (wires) extending therethrough that electrically connect the electrode contacts 126 to the stimulator unit 120.

[0031] Figure 1A Also shown is an extracochlear electrode contact 126 (23). The extracochlear electrode contact 126 (23) is an electrical contact that is configured to, for example, deliver electrical stimulation to a recipient's cochlea and / or sink electrical current from a recipient's cochlea. The extracochlear electrode contact 126 (23) is connected to a reference lead 123 that includes one or more conductors that electrically couple the extracochlear electrode contact 126 (23) to the stimulator unit 120.

[0032] Typically, the electrode contacts of commercial cochlear implants and many other types of prostheses that stimulate tissue are formed of biocompatible and stable metallic materials (e.g., metals or metal alloys), such as platinum group metals (e.g., platinum, osmium, iridium, ruthenium, rhodium, and palladium), alloys of platinum with a small percentage of iridium, alloys of stainless steel, titanium, gold, etc. Although very stable, these metallic materials do have a finite dissolution rate, especially under conditions of high charge stimulation. Ex vivo analysis of chronically implanted cochlear implants has shown that some electrode contacts show signs of dissolution over extended periods of time (e.g., decades).

[0033] Typically, electrode contacts formed from metallic materials are formed from atoms from which electrons are separated and free to move in the solid. As used herein, "dissolution" of an electrode contact refers to the metal losing electrons to form metal ions (Me) that can move out of the solid. v+The process of metal dissolution is a kinetic process and is quantified by its rate. In prostheses that stimulate tissue, metal dissolution is caused by the fact that the electrode contacts are often surrounded by or in contact with body fluids that lead to electron loss and metal ion detachment. This process is often accelerated in stimulating prostheses because the potential of the metal changes during the delivery of the stimulation pulse. This may accelerate the dissolution process for two reasons: i) metal dissolution is often thermodynamically favored at higher metal potentials, and ii) the changing potential leads to a rearrangement of surface metal atoms (e.g. due to oxide formation and reduction), making it more likely that metal atoms will ionize and move away from the metal surface. The consequence of the removal of metal ions is a reduction in the mass of the electrode contacts (i.e., electrode contact corrosion).

[0034] It is expected that if dissolution proceeds indefinitely, dissolution may affect the integrity of the electrode contacts (e.g., eventually a large portion of the material may be removed from the electrode contacts, causing the contacts to become more high impedance and eventually cease to conduct current). Thus, dissolution may limit the useful life of cochlear implants and other prostheses that stimulate tissue for at least a subset of the implanted population. Dissolution may also affect future technologies such as thin-film electrodes and multipolar stimulation.

[0035] Techniques are presented herein for improving the long-term efficacy of a prosthesis for stimulating tissue by minimizing the dissolution rate of the tissue stimulator electrode contacts. More specifically, according to embodiments presented herein, a continuous dissolving barrier surrounds / encapsulates a stimulation assembly, the latter including the electrode contacts and the carrier member. The dissolving barrier is formed of a material that allows the conduction of electrical current, primarily through a chlorine (e.g., Cl - ) and sodium (e.g., Na + ) in the tissue, but inhibits the movement of metal ions from the surface of the electrode contact. There are many characteristics that distinguish electrode metal ions from major carrier ions in tissue. For example, electrode metal ions are usually heavy (e.g., the atomic mass of platinum is 195, while the atomic mass of chlorine is 35 and the atomic mass of sodium is 23). In addition, electrode metal ions are usually multivalent (e.g., platinum has two stable ionic forms (Pt 2+ and Pt 4+ ), while the sodium and chloride ions in the solution are monovalent (Cl + 、Na -). In addition, heavy metal ions such as platinum are often chelated (i.e., form stable complexes) with organic molecules such as proteins, peptides, and amino acids, while sodium and chloride ions are freely present in large quantities in the tissue environment. All of these physical differences between electrode metal ions and most carrier ions in tissue can be used to selectively block or reduce the movement of electrode metal ions while still allowing the movement of carrier ions. Other possible types of barrier materials rely on the non-ionic movement of charge to inhibit the movement of electrode metal ions. For example, the movement of charge through conductive polymers is generated via the movement of so-called "polarons," which exchange charge with ions on the surface of the material. Ions themselves cannot move directly through conductive polymers, so such materials can effectively block the movement of all ions from the metal surface while allowing charge to be conducted through themselves. Other material types, such as, for example, carbon nanotubes, graphene, and conductive diamonds also conduct charge in a non-ionic manner and can therefore be used to conduct current and simultaneously inhibit electrode metal dissolution.

[0036] exist Figure 1B , a continuous dissolving barrier is generally illustrated by reference numeral 150. As noted, dissolving barrier 150 reduces electrode contact dissolution, thereby extending the working life of the in situ cochlear implant 100. Further details regarding a continuous dissolving barrier, such as barrier 150, are provided below.

[0037] In particular, Figure 2A is a cross-sectional view of a continuous dissolving barrier 250 according to certain embodiments presented herein. The continuous dissolving barrier 250 encapsulates a stimulation assembly 218 that includes an elongated insulating carrier member 234 and a plurality of spaced-apart electrode contacts 226 disposed along the length of the carrier member. Figure 2A A side view of the carrier member 234 and the plurality of electrode contacts 226 is shown, while the dissolving barrier 250 is shown in cross-section for ease of illustration. In practice, the dissolving barrier 250 is disposed substantially over the entire surface of the stimulation component 218 (i.e., the dissolving barrier 250 substantially encapsulates the entire carrier member 234 and the electronic contacts 226).

[0038] Figure 2B yes Figure 2A , which shows a segment of the dissolution barrier 250 and the electrode contact 226. As shown by arrow 252, the dissolution barrier 250 is configured so that current flows through the dissolution barrier 250. However, the dissolution barrier 250 is also configured so that the dissolution barrier 250 blocks metal (e.g., platinum) atoms and ions, as described above. That is, the platinum atoms and ions remain trapped behind (below) the dissolution barrier 250 (i.e., the platinum atoms and ions remain at the electrode contact 226).

[0039] According to embodiments presented herein, the dissolving barrier 250 can be a continuous coating, sheath, sheath, membrane, layer, film, or other type of barrier applied to the carrier member 234 and the electrode contacts 226. The dissolving barrier 250 is configured to inhibit the rate of removal of metal (e.g., platinum) ions from the electrode contact surfaces and has a sufficiently high conductivity to allow current to pass through / through it without adding excessive impedance to the neural stimulation path (i.e., without adding excessive impedance to the current signal generated or absorbed by the electrode contacts). However, the dissolving barrier 250 is configured so that it has a sufficiently low conductivity to inhibit significant shunting of current between the electrode contacts.

[0040] Thus, as further described below, a dissolution barrier according to embodiments presented herein, such as dissolution barrier 250, has a conductivity below an upper limit / threshold, but above a lower limit / threshold. If the conductivity of the dissolution barrier is too high, then the dissolution barrier will shunt too much current between the electrode contacts (e.g., as compared to a layer that is only on the electrode contacts and does not have an upper conductivity limit).

[0041] Furthermore, if the conductivity of the dissolving barrier is too low, then the dissolving barrier will significantly increase the impedance of the electrode, consuming more power and limiting the current flow to the target nerve (e.g., the auditory nerve). Thus, either too high or too low conductivity can render the tissue stimulating prosthesis unusable for its intended purpose and / or negatively impact the performance of the tissue stimulating prosthesis. The appropriate conductivity range for the dissolving barrier is discussed below and depends on the thickness of the barrier material and the size of the electrode contacts.

[0042] It should be noted that reducing the thickness of the barrier material may be beneficial in two ways: reducing the added impedance and reducing the shunt current flowing therein. Therefore, in general, the barrier material should be as thin as possible. However, limitations of the application process, such as the thickness of the dip coating or the brittleness of the applied film, may dictate the minimum thickness of the barrier material. Once this minimum thickness is determined (determined based on the application process), the conductivity of the applied barrier material can be adjusted to ensure that the added impedance and shunt current are within tolerable limits. The conductivity of the candidate barrier material can often be adjusted by adjusting the chemistry of the manufacturing process. For example, the conductivity of a conductive polymer can be adjusted by a method called organic synthesis, in which different organic molecules are incorporated into the molecular structure of the polymer, thereby affecting its conductivity. Typically, even a very thin layer of barrier material is sufficient to block the flow of electrode metal ions. Therefore, when determining the thickness of the barrier material, the efficiency of the barrier material in blocking metal electrode ions is rarely a factor.

[0043] Figure 3A 1 generally illustrates the variables used in calculating different parameters of the dissolution barrier according to the embodiments presented herein. More specifically, Figure 3A An electrode contact 366 having a surface area A and a portion of a dissolution barrier 350 made of a material having a conductivity C and a thickness T, where the conductivity is measured in Siemens per meter (Sm -1 The impedance Z of the dissolving barrier 350 is then given by the following equation 1.

[0044] Equation 1:

[0045] Z=T / (AC)

[0046] Equation 1 can be arranged to solve for C as shown in Equation 2 below.

[0047] Equation 2:

[0048] C=T / (AZ)

[0049] If the impedance Z has a maximum acceptable value Zmax, the formula for the minimum conductivity Cmin is as shown in Equation 3 below.

[0050] Equation 3:

[0051] Cmin=T / (A.Zmax)

[0052] In one example, assume that the surface area of ​​each of the electrode contacts 326 is 0.15 mm 2 (ie, A is equal to 0.15 mm 2 , which is typical for some cochlear implant electrode contacts). It is also assumed that the dissolving barrier 350 has a thickness of 0.1 mm and is made of a material having a conductivity C Sm-1 (i.e., T equals 0.1 mm and C is an unknown variable). It is also assumed that the acceptable maximum impedance of the dissolving barrier 350 is 1 kilo-ohm (kΩ) (i.e., Zmax equals 1 kΩ). In this example, the example acceptable maximum impedance is based on the total impedance of the electrode contacts of the above dimensions in a typical cochlea. In particular, 1 kΩ is approximately 10% of the total impedance of electrodes of these dimensions in a typical cochlea.

[0053] If A = 0.15 mm is used in Equation 3 2 , T = 0.1 mm and Zmax 1 kΩ, the minimum value of C (i.e. Cmin) is about 0.67 Sm-1 (i.e., 10 -4 / (1.510 -7 .10 3 )). This is the minimum acceptable value of the conductivity C of the dissolution barrier 350, as any higher value would result in an impedance of the dissolution barrier greater than 1 kΩ. If the maximum acceptable impedance value changes, the minimum acceptable value of the conductivity will also change. It should be noted that all units in the above equations are expressed in SI units.

[0054] As described above, in a cochlear implant, the stimulation assembly is implanted in one of the recipient's cochlear ducts / ear canals, i.e., in the scala vestibuli or scala tympani. Each of the scala vestibuli and scala tympani is filled with a conductive fluid called perilymph. The perilymph of the scala vestibuli and scala tympani has a composition similar to cerebrospinal fluid (CSF), i.e., rich in sodium (140 mM) and lacking in potassium (5 mM) and calcium (1.2 mM). The perilymph of the scala vestibuli comes from plasma that crosses the hemto-perilymphatic barrier, while the perilymph of the scala tympani comes from CSF. The perilymph is conductive to electrical current.

[0055] As noted, in the above example, the minimum conductivity of the dissolution barrier is approximately 0.67 Sm -1 Now, assume that the maximum acceptable conductivity of the barrier material is the same as the conductivity of the perilymph, which is about 180 Sm -1 , which is similar to plasma. If the conductivity is higher than that of the perilymph, then this will increase the current flow between the electrode contacts during stimulation.

[0056] Figure 3B is a schematic diagram illustrating the current shunting between electrode contacts if the conductivity of the dissolution barrier 350 is set too high. Figure 3B , two electrode contacts 326 are disposed in an insulating carrier member 334 (e.g., formed of silicone) and are separated by a segment / portion 355 of the carrier member 334 having a length D (i.e., the electrode contacts are separated / spaced apart from each other by a distance D). As indicated by arrow 354, too high an electrical conductivity of the dissolving barrier 350 minimizes a normal current path into the recipient tissue. As indicated by arrow 356, the minimization of normal current flow results in excessive current flow between the electrode contacts 326 through / via the dissolving barrier 350.

[0057] Note that the conductivity of the perilymph is quite high and causes significant power losses in conventional cochlear implants. Selecting the same upper conductivity limit as the conductivity of the perilymph means that the dissolving barrier does not make the electrical situation worse (i.e., no additional current is shunted between the electrode contacts due to the presence of the dissolving barrier). In some embodiments, the conductivity of the dissolving barrier can be less than the conductivity of the perilymph or other surrounding body tissue / fluid. For example, the conductivity of the dissolving barrier can be about 5% or 10% less than the conductivity of the perilymph. However, in some embodiments, if the tradeoff (e.g., better dissolution prevention) is sufficiently high, it may be acceptable to use a dissolving barrier with a conductivity slightly higher than the conductivity of the perilymph. Practical disadvantages of using a higher conductivity material may be higher power consumption and possible increased diffusion of excitation of the auditory nerve.

[0058] In the above example, for a material having a thickness T of approximately 0.1 mm and a thickness T of 0.15 mm2 The acceptable conductivity range of the dissolution barrier for the electrode area A is approximately 0.67 Sm -1 and about 180Sm -1 Between (0.67Sm -1 <C<180Sm -1 ). It will be appreciated that for other barrier thicknesses this range will be different. Again assuming a maximum additional impedance of 1 kΩ and an electrode area of ​​1.5.10 - 7 m 2 (0.15mm 2 ), then, as shown in Equation 4 below, Equation 3 can be written to find the minimum acceptable conductivity Cmin based on the thickness T of the barrier material.

[0059] Equation 4:

[0060] Cmin=T / 1.5.10-4

[0061] Therefore, for a thickness T equal to 0.05 mm, the minimum conductivity C min is 5.10 -5 / 1.5.10 -4 , or 0.33 Sm-1. Therefore, for a 0.05 mm thick dissolution barrier, the conductivity range C is between approximately 0.33 Sm -1 and about 180Sm -1 Between (0.33 <C<180Sm -1 ).

[0062] As described above, the dissolving barrier according to the embodiments presented herein has a conductivity below an upper limit / threshold, but also has a conductivity above a lower limit / threshold. If the conductivity of the dissolving barrier is too high, then the dissolving barrier will shunt too much current between the electrode contacts, and if the conductivity is too low, it will increase the impedance to the stimulation path. In other words, the dissolving barrier has both sufficient conductivity without adding significant impedance to the electrical stimulation path (e.g., more than 1 kOhm), and sufficient electrical insulation without shunting significant current between the electrode contacts. In addition to the above, the dissolving barrier according to the embodiments presented herein may also have additional properties.

[0063] For example, a dissolving barrier according to embodiments presented herein can be configured to be applied via a reproducible manufacturing process. Examples of reproducible manufacturing processes include, for example, dip coating, spray coating, shrink wrapping, and the like. In addition, a dissolving barrier according to embodiments presented herein is biocompatible, does not degrade during sterilization, and remains stable in the cochlea for decades. Furthermore, a dissolving barrier according to embodiments presented herein is configured such that it does not affect the mechanical properties of the stimulation assembly and is close to or in contact with the carrier member and electrode contacts (e.g., to avoid increasing the cross-sectional area of ​​the array, making it more difficult to insert). The dissolving barrier is typically relatively thin in order to: minimize additional stimulation impedance for a given material conductivity, minimize current shunting between adjacent electrode pads for a given material conductivity, minimize the total cross-section of the electrode array, and minimize any mechanical changes to the array caused by the mechanical properties of the barrier.

[0064] In a specific example, dip coating is used to apply a dissolving barrier to the stimulation component. For example, at the end of the stimulation component manufacturing process, the completed stimulation component (i.e., the electrode embedded in the carrier member) is immersed in a solution containing a soluble form of the barrier material. Once the material dries or hardens, it forms a strong, flexible coating on the stimulation component. In another example, the liquid barrier material is applied via a device that sprays a mist onto the stimulation component. In other examples, the shrink wrap process takes a thin, solid form of the barrier material (e.g., similar to plastic wrap or a blister pack) and wraps it around the stimulation component. Once the wrap is in place, heat is applied by a heat gun or similar device, and the wrap will shrink to fit snugly on the stimulation component.

[0065] The dissolution barrier according to the embodiments presented herein can be formed from a variety of different material types that meet many of the above-mentioned characteristics. For example, in certain embodiments, the dissolution barrier can be formed from a biocompatible conductive polymer or a biocompatible conductive hydrogel. In further embodiments, the dissolution barrier can be formed from anti-dissolution nanoparticles such as carbon nanotubes suspended in a biocompatible polymer, plastic, silicone or other insulating solid. In further embodiments, the dissolution barrier can be formed from a woven biocompatible material, such as an insulating polymer or plastic. For example, a fine woven mat of insulating material can be made with sufficient porosity to allow charge carrying ions to pass through while reducing the flow of larger metal ions, such as platinum ions.

[0066] In certain embodiments, the dissolution barrier may be formed by a self-assembled monolayer. Figure 4As shown, the self-assembled monolayer 462 is formed of carbon chain molecules 462 having a thiol or "head" group 464 that promotes adhesion to a conductor and having a functional group 466 at its other end, such as a protein. The thiol group 464 can be used to firmly attach the carbon chain molecules 462 to a surface that may not be easily attached without the self-assembled monolayer 462. If the carbon chains 462 are made shorter, the self-assembled monolayer 462 can present a low resistance to the passage of current, and depending on the functional group, it can also effectively inhibit the diffusion of metal ions from the surface.

[0067] In some examples, the dissolution barrier may be formed by an ion permeable membrane. An ion permeable membrane is a material that allows the movement of conductive ions through drift, but blocks the diffusion of platinum ions. This selectivity (for conductive ions through drift rather than for diffusion of platinum ions) can be based on several factors. For example, the selectivity can be based on the size of the ions, where platinum ions (e.g., [PtCl4] 2- ) is large, while the main carrier ions (e.g., Cl - , Na+, K+) are smaller. Selectivity can also or alternatively be based on the valence of the ions (i.e., platinum ions are divalent and carrier ions are monovalent). Selectivity can also or alternatively be based on actual drift vs. diffusion characteristics, where carrier ions move under the influence of the electric field, but diffusing ions do not. Selectivity can also or alternatively be based on frequency, where the load ions move quickly in the stimulating tissue prosthesis, while the platinum ions move much slower under the influence of diffusion and other transport mechanisms in the body.

[0068] In some examples, the dissolution barrier can be formed from doped crystalline materials, such as nitrogen-doped diamond and titanium nitride. These materials can provide a low-resistance interface and are typically applied using sputtering, vapor or other deposition techniques.

[0069] Embodiments are described herein primarily with reference to dissolution barriers that are primarily configured to inhibit dissolution of electrodes without negatively impacting stimulation pathways and without burdening the manufacturing process. That is, dissolution barriers according to embodiments presented herein are both sufficiently strong to prevent the transport of platinum atoms, ions, or particles (i.e., substantially prevent dissolution of platinum electrode contacts in the presence of body fluids) and have a conductivity above a lower limit / threshold, but still below an upper limit / threshold (i.e., an appropriate conductivity to allow stimulation current to pass through the barrier without significantly increasing impedance and without shorting adjacent electrode contacts to each other). In addition to the foregoing, dissolution barriers according to embodiments presented herein may also have additional properties for other functions.

[0070] For example, in addition to the platinum blocking and conductivity limiting described above, dissolution barriers according to the presented embodiments may be configured for drug elution, lubrication, and the like.

[0071] As described above, the electrode contacts are used to deliver stimulation signals (electric current) to the recipient's tissue. In some prostheses, each of the electrode contacts is a solid conductive element of a certain thickness (e.g., cut or stamped from a sheet or tube of platinum, a platinum alloy, etc.). Recently, attempts have been made to use microfabrication techniques, such as photolithography, thin film deposition, etc., to manufacture electrode contacts on, for example, a silicon substrate. For example, thin film deposition is a process in which a very thin film of material (e.g., a few nanometers to about 100 microns) is applied to a "substrate" surface or a previously deposited coating. Therefore, a "thin film" or "micro" electrode contact is composed of several layers of thin conductive material stacked on top of each other on an insulating substrate.

[0072] The final thickness of the thin film electrode may vary depending on, for example, the application, the manufacturing process, etc. However, thin film electrodes are typically much thinner than conventional electrode contacts and, therefore, are susceptible to long-term problems due to dissolution. That is, because the thickness of the thin film electrode is less than that of a conventional electrode, the thin film electrode can withstand less material loss before failure. Therefore, the dissolution barrier according to the embodiments presented herein may be particularly advantageous in thin film electrodes.

[0073] In one example, a preformed solid film of barrier material can be applied to a thin film electrode, since the film only needs to coat one side of the electrode, which is essentially both sides. A separate film of barrier material can be stretched over the electrode and trimmed away. A thin layer of adhesive can be applied to the film to hold it in place and / or electrostatic forces can be used to hold the film in place (e.g., as with plastic wrap or blister packaging). In some embodiments, the barrier material can be added as an additional step in the film processing.

[0074] The embodiments presented herein have been primarily described with reference to cochlear implants. However, as noted elsewhere, the techniques presented herein may also or alternatively be used with other types of prostheses that stimulate tissue (e.g., auditory brainstem stimulators, implantable pacemakers, spinal cord stimulators, deep brain stimulators, motor cortex stimulators, sacral nerve stimulators, pudendal nerve stimulators, vagus nerve / vagal nerve stimulators, trigeminal nerve stimulators, retinal or other visual prostheses / stimulators, occipital cortex implants, diaphragm (phrenic) pacemakers, analgesic stimulators, other nerves or neuromuscular stimulators, etc.). Figure 5 is a simplified schematic diagram illustrating an exemplary spinal cord stimulator 500 in which the embodiments presented herein may be implemented.

[0075] The spinal cord stimulator 500 includes an implant (main module) 514, a lead region 516, and a stimulation assembly 518. The implant 514 includes a stimulator unit 520, which includes, among other things, one or more current sources on an integrated circuit (IC).

[0076] The stimulation assembly 518 is implanted in the recipient adjacent to / near the recipient's spinal cord 527 and includes five (5) stimulation electrodes 526, referred to as stimulation electrodes 526(1)-526(5). The stimulation electrodes 526(1)-526(5) are disposed in an electrically insulating carrier member 534 and are electrically connected to the stimulator 520 via conductors (not shown) extending through the carrier member 534.

[0077] After implantation, stimulator unit 520 generates stimulation signals for delivery to spinal cord 527 via stimulation electrodes 526(1)-526(5). Figure 5 Not shown, but an external controller may also be provided to transmit signals through the recipient's skin / tissue to the stimulation electronics 633 to control the stimulation signals.

[0078] Similar to the above-described embodiments, the stimulation assembly 518 includes a dissolving barrier 550 configured to reduce electrode contact dissolution, thereby extending the operating life of the in situ spinal cord stimulator 500. The dissolving barrier 550 is configured to allow current to flow through / through the dissolving barrier 550, but is also configured to allow the dissolving barrier 550 to block platinum atoms and ions. That is, the platinum atoms and ions remain trapped behind (below) the dissolving barrier 5250 (i.e., the platinum atoms and ions remain at the electrode contacts 526).

[0079] According to embodiments presented herein, the dissolving barrier 550 can be a continuous coating, sheath, sheath, membrane, layer, film, or other type of barrier applied to the carrier member 534 and the electrode contacts 526. The dissolving barrier 550 is configured to reduce the rate of removal of platinum ions from the electrode contact surfaces and has a sufficiently high conductivity to allow current to pass through / through it without adding excessive impedance to the neural stimulation path. However, the dissolving barrier 550 is configured so that it has a sufficiently low conductivity to prevent significant shunting of current between the electrode contacts.

[0080] Figure 6 660 is a flow chart of a method 660 according to certain embodiments presented herein. The method 660 begins at 662, wherein a plurality of electrode contacts arranged in an elongated array are provided. At 664, the plurality of electrode contacts are encapsulated in an implantable carrier member formed of an insulating material. The plurality of electrode contacts extend along the longitudinal axis of the carrier member and are positioned at an outer surface of the carrier member. At 666, a surface of each of the plurality of electrode contacts is exposed for applying an electrical stimulation signal to adjacent neural tissue. At 668, a continuous dissolution barrier is laminated on each of the plurality of electrode contacts and the carrier member.

[0081] It should be understood that the above-described embodiments are not mutually exclusive and that the various embodiments may be combined in various ways and arrangements.

[0082] The scope of the present invention described and claimed herein is not limited by the specific preferred embodiments disclosed herein, because these embodiments are intended to be illustrative of several aspects of the present invention and not limiting. Any equivalent embodiments are intended to be within the scope of the present invention. In fact, from the foregoing description, in addition to those shown and described herein, various modifications to the present invention will become apparent to those skilled in the art. Such modifications are also intended to fall within the scope of the appended claims.

Claims

1. A prosthesis for stimulating tissue, comprising: a stimulation assembly including an elongated carrier member and a plurality of spaced apart electrode contacts disposed along the elongated carrier member, the plurality of electrode contacts configured to perform at least one of providing a current signal or sinking a current signal; and a dissolving barrier encapsulating the stimulation component, wherein the dissolving barrier is configured to inhibit in situ dissolution of the plurality of electrode contacts; Wherein the dissolution barrier is configured to allow electrical current conduction but inhibit movement of multivalent metal ions from surfaces of the plurality of electrode contacts.

2. The tissue stimulating prosthesis of claim 1, wherein the dissolving barrier has an electrical conductivity, and wherein the electrical conductivity is greater than a first threshold and less than a second threshold.

3. The tissue stimulating prosthesis of claim 2, wherein the first threshold and the second threshold are selected based on a surface area of ​​the plurality of electrode contacts and a thickness of the dissolving barrier.

4. The tissue stimulating prosthesis of claim 2, wherein the stimulation component is configured to be implanted in the recipient at a location where the dissolving barrier contacts a bodily fluid of the recipient, and wherein the second threshold is the same as the conductivity of the bodily fluid.

5. The tissue stimulating prosthesis of claim 2, wherein the stimulation component is configured to be implanted in the recipient at a location where the dissolving barrier contacts a bodily fluid of the recipient, and wherein the second threshold is less than the conductivity of the bodily fluid.

6. The tissue stimulating prosthesis of claim 5, wherein the stimulating assembly is configured to be implanted in the cochlea of ​​a recipient such that the dissolving barrier contacts the recipient's perilymph, and wherein the second threshold is 180 siemens per meter (Sm -1 ).

7. The tissue stimulating prosthesis of claim 2, wherein the first threshold is a level at which the impedance of the dissolving barrier is less than 10% of the total impedance of each of the plurality of electrode contacts.

8. The tissue stimulating prosthesis of claim 2, wherein the acceptable maximum impedance of the dissolving barrier is 1 kilo-ohm (kΩ).

9. The tissue stimulating prosthesis according to claim 2, wherein the first threshold value is a minimum conductivity C given in Siemens per meter. min , and where: C min =T / 1.5·10 -4 , where T is the thickness of the barrier material in meters.

10. The tissue stimulating prosthesis of claim 1, wherein the dissolving barrier comprises a biocompatible conductive polymer.

11. The tissue stimulating prosthesis of claim 1 , wherein the dissolving barrier comprises a biocompatible conductive hydrogel.

12. The tissue stimulating prosthesis of claim 1, wherein the dissolution barrier comprises dissolution-resistant nanoparticles suspended in at least an insulating solid.

13. The tissue stimulating prosthesis of claim 1, wherein the dissolving barrier comprises a biocompatible braided insulating polymer.

14. The tissue stimulating prosthesis of claim 1, wherein the dissolving barrier comprises plastic.

15. The tissue stimulating prosthesis of claim 1, wherein the dissolving barrier comprises a self-assembled monolayer.

16. The tissue stimulating prosthesis of claim 1, wherein the dissolution barrier comprises an ion-permeable membrane configured to selectively allow movement of conductive ions by drift while preventing diffusion of metal ions.

17. The tissue stimulating prosthesis of claim 16, wherein the ion-permeable membrane is configured to selectively allow conductive ions to move by drift while preventing diffusion of metal ions based on at least one of the size of the ions, the valence of the ions, or the frequency of movement of the ions.

18. The tissue stimulating prosthesis of claim 1, wherein the plurality of electrode contacts are thin film electrode contacts.

19. A device for stimulating tissue, comprising: an implantable carrier member formed of an insulating material and having a central longitudinal axis and an outer surface; a plurality of electrode contacts distributed along the longitudinal axis on the outer surface of the carrier member for applying electrical stimulation signals to adjacent neural tissue; and a continuous dissolution barrier laminated to each of the plurality of electrode contacts and the carrier member; Wherein the dissolution barrier is configured to allow electrical current conduction but inhibit movement of multivalent metal ions from surfaces of the plurality of electrode contacts.

20. The device of claim 19, wherein the plurality of electrode contacts are each spaced apart from one another by segments of the carrier member, and wherein the continuous dissolution barrier is laminated over the plurality of electrode contacts and over the segments of the carrier member between the plurality of electrode contacts.

21. The device of claim 19, wherein the dissolution barrier has an electrical conductivity, and wherein the electrical conductivity is greater than a first threshold and less than a second threshold.

22. The apparatus of claim 21, wherein the first threshold and the second threshold are selected based on a surface area of ​​the plurality of electrode contacts and a thickness of the dissolution barrier.

23. The device of claim 21, wherein the device is configured to be at least partially implanted in the recipient at a location where the dissolving barrier contacts a bodily fluid of the recipient, and wherein the second threshold is the same as the conductivity of the bodily fluid.

24. The device of claim 21, wherein the device is configured to be at least partially implanted in the recipient at a location where the dissolving barrier contacts a bodily fluid of the recipient, and wherein the second threshold is less than the conductivity of the bodily fluid.

25. The device of claim 21, wherein the device is configured to be at least partially implanted in the cochlea of ​​a recipient such that the dissolving barrier is in contact with the recipient's perilymph, and wherein the second threshold is 180 Siemens per meter (Sm -1 ).

26. The device of claim 21, wherein the first threshold is a level at which the impedance of the dissolving barrier is less than 10% of the total impedance of each of the plurality of electrode contacts.

27. The device of claim 21, wherein the acceptable maximum impedance of the dissolving barrier is 1 kiloohm (kΩ).

28. The apparatus of claim 21, wherein the first threshold is a minimum conductivity C given in Siemens per meter. min , and where: C min =T / 1.5·10 -4 , where T is the thickness of the barrier material in meters.

29. The device of claim 19, wherein the dissolving barrier comprises a biocompatible conductive polymer.

30. The device of claim 19, wherein the dissolving barrier comprises a biocompatible conductive hydrogel.

31. The device of claim 19, wherein the dissolution barrier comprises dissolution-resistant nanoparticles suspended in at least an insulating solid.

32. The device of claim 19, wherein the dissolving barrier comprises a biocompatible braided insulating polymer.

33. The device of claim 19, wherein the dissolving barrier comprises plastic.

34. The device of claim 19, wherein the dissolution barrier comprises a self-assembled monolayer.

35. The apparatus of claim 19, wherein the dissolution barrier comprises an ion-permeable membrane configured to selectively allow movement of conductive ions through drift while preventing diffusion of metal ions.

36. The apparatus of claim 35, wherein the ion-permeable membrane is configured to selectively allow movement of conductive ions by drift while preventing diffusion of metal ions based on at least one of the size of the ions, the valence of the ions, or the frequency of movement of the ions.

37. The device of claim 19, wherein the plurality of electrode contacts are thin film electrode contacts.

38. A method for providing a plurality of electrode contacts, comprising: providing a plurality of electrode contacts arranged in an elongated array; encapsulating the plurality of electrode contacts in an implantable carrier member formed of an insulating material, wherein the plurality of electrode contacts extend along a longitudinal axis of the carrier member and are positioned at an outer surface of the carrier member; exposing a surface of each of the plurality of electrode contacts for applying an electrical stimulation signal to adjacent neural tissue; and laminating a continuous dissolution barrier over each of the plurality of electrode contacts and the carrier member; Wherein the dissolution barrier is configured to allow electrical current conduction but inhibit movement of multivalent metal ions from surfaces of the plurality of electrode contacts.

39. The method of claim 38, wherein laminating a continuous dissolution barrier over each electrode contact of the plurality of electrode contacts and the carrier member comprises: The carrier member having electrode contacts is dip coated.

40. The method of claim 38, wherein laminating a continuous dissolution barrier over each of the plurality of electrode contacts and the carrier member comprises: The continuous dissolving barrier is sprayed onto each of the plurality of electrode contacts and the carrier member.

41. The method of claim 38, wherein the continuous dissolution barrier has an electrical conductivity, and wherein the electrical conductivity is greater than a first threshold and less than a second threshold.

42. The method of claim 41, wherein the acceptable maximum impedance of the dissolving barrier is 1 kiloohm (kΩ).

43. A method according to claim 41 or 42, wherein the first threshold value is a minimum conductivity Cmin given in Siemens per meter, and wherein: Cmin=T / 1.5·10 -4 , where T is the thickness of the barrier material in meters.

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