US 20080033520 A1
An implantable electrode with a fluid reservoir is described. An implantable electrode carrier has an outer surface including electrode contacts for electrically stimulating nearby neural tissue. A fluid reservoir within the electrode carrier stores a volume of therapeutic fluid. At least one fluid delivery port connects the fluid reservoir to the outer surface of the electrode carrier for delivering the therapeutic fluid from the fluid reservoir to the outer surface. A delivery septum port is in fluid communication with the fluid reservoir for delivering the therapeutic fluid to the fluid reservoir.
1. A cochlear implant electrode array comprising:
a cochlear electrode array for electrically stimulating cochlear tissue, the array including a drug eluting portion having a pharmaceutical agent incorporated into a polymer material,
wherein the drug eluting portion is adapted to release over time a therapeutically effective amount of the pharmaceutical agent for the inner ear.
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The present application is a divisional of U.S. application Ser. No. 11/492,212, filed Jul. 24, 2006, which claimed priority from U.S. Provisional Application Ser. No. 60/780,667, filed Mar. 9, 2006, and which also was a continuation-in-part of U.S. application Ser. No. 11/374,505, filed Mar. 13, 2006, which was a divisional application of U.S. application Ser. No. 10/281,066, filed Oct. 24, 2002, and issued as U.S. Pat. No. 7,044,942, which in turn claimed priority from U.S. Provisional Application Ser. No. 60/336,452, filed Oct. 24, 2001; U.S. Provisional Application Ser. No. 60/394,427, filed Jul. 8, 2002; U.S. Provisional Application Ser. No. 60/394,602, filed Jul. 9, 2002; and U.S. Provisional Application Ser. No. 60/417,704, filed Oct. 10, 2002; all of which are hereby incorporated by reference.
The invention relates to a drug eluting cochlear implant electrode for the transient elution of a pharmacological agent into the inner ear.
Electrical stimulation of the inner ear has been very successful in restoring sound sensation to patients afflicted with deafness. Intra-cochlear electrodes are intended to restore some sense of hearing by direct electrical stimulation of the neural tissue in proximity of an electrode contact. The electrical stimulation is accomplished with an implanted cochlear implant stimulator connected to an electrode inserted deep into the scala tympani cavity.
But the insertion of the electrode causes a variable amount of connective tissue growth and trauma. The amount of trauma is very difficult to predict and depends on the cochlea anatomy, the electrode design, and the insertion technique. The trauma inflicted to the tissues may subsequently cause apoptosis and/or necrosis of nervous tissue (i.e., hair cells and spiral ganglion cells). Tissue growth and trauma may limit the performance of the implant. And trauma to spiral ganglion cells is cumulative and cannot be undone in the present state of technology. As more patients with significant usable residual hearing receive a cochlear implant, it becomes ever more important to use a minimally traumatic electrode. And as more patients are implanted at a young age who will be re-implanted several times during their lifetime, each consecutive insertion should limit the trauma to spiral ganglion cells to a minimum.
Trauma is usually caused by the electrode insertion into the delicate tissue of the inner ear. Insertion requires mechanical forces to be applied on the electrode to overcome the friction of the electrode against the tissue of the spiraling cochlea. To reduce trauma to the organ or tissue, electrodes and catheters should be soft and flexible, and insertion forces should be minimum. Unfortunately, most cochlear implant electrodes on the market today require significant force to be inserted, even for distances which are much less than the full length of the scala tympani.
The force required to insert an electrode or catheter is related to its size, geometry, and fabrication material. Material used in such devices includes materials for wires, contacts, functional metallic or polymer segment, and bulk material. The size of the device, the rigidity of the material used, the hydrophobicity of the outer shell of the electrode array, the energy stored in one way or another in the electrode, and the insertion process of the device all have an impact on the amount and location of tissue damage that will be inflicted during electrode placement.
Damage and trauma cause bleeding, inflammation, perforation of soft tissue, tears and holes in membranes, and fracture of thin osseous structures. The resulting damage may cause loss of surviving hair cells, retrograde degeneration of the dendrite which inervates the organ of Corti, and in the worst case, spiral ganglion cell death in the Rosenthal's canal. Cell death means that quantitatively less neural tissue is available for stimulation, and qualitatively that fewer frequency-tuned fibers are available to represent frequency information. Further loss of hair cells and loss of dendrites without loss of spiral ganglion cells means that acoustic stimulation is no longer possible, and that no synergetic effects between acoustic and electric stimulation is available. Electro-acoustic synergetic effects may be important for good sound discrimination in noisy environments.
Another inconvenience with cochlear implants is the rise in measured electrode impedance post-surgery. This rise is thought to be caused by encapsulation of the electrode by a tight membrane which reduces the efficiency of electric stimulation by creating a zone with ionic depletion around the contacts. It would make sense to post-surgically introduce some medicine into the cochlea to maintain a lower electrode impedance. It has been demonstrated, for example, that the introduction of cortico steroids can reduce the impedance rise after surgery. This has been done by depositing or rubbing the medicine on the electrode. But as the electrode is introduced in the fluid of the scala tympani, the medical solution quickly dissolves and may not reach a location where it would be most beneficial.
There have been attempts with non-cochlear implant patients to deliver medicine to the inner ear for the treatment of Meniere's disease or vertigo. The drug delivery takes place through the somewhat permeable round window membrane after injection of a bolus into the middle ear. One problem with round window drug delivery is that the membrane permeability to molecular substances changes over the course of a day, and that large molecules cannot pass through the tight membrane. It is thought that very little pharmacologic substance reaches the cochlear region beyond the first few millimeters of cochlea length.
There is no easy existing way to deliver medicine into the inner ear after cochlear implantation. The middle ear is not easily accessed and the inner ear is a sealed system that does not allow direct deposition or injection of medicines except at the time of cochlear implant surgery. After surgery the cochlea is partially filled with the electrode which should not be moved or removed.
Drug eluting electrode leads with cortico steroids have been used successfully in the past with cardiac pacemaker electrodes to reduce the contact impedance. In addition, silicone elastomer loaded with a pharmacological agent has been used as an eluting structure in several applications such as birth control, vascular injury treatment, and stents. Drug eluting electrodes have not been used with cochlear implants.
Embodiments of the present invention include an implantable electrode with a fluid reservoir. An implantable electrode carrier has an outer surface including electrode contacts for electrically stimulating nearby neural tissue. A fluid reservoir within the electrode carrier stores a volume of therapeutic fluid. At least one fluid delivery port connects the fluid reservoir to the outer surface of the electrode carrier for delivering the therapeutic fluid from the fluid reservoir to the outer surface. A delivery septum port is in fluid communication with the fluid reservoir for delivering the therapeutic fluid to the fluid reservoir.
In further embodiments, a fluid delivery channel connects the delivery septum port to the fluid reservoir. Embodiments may also include a return septum port in fluid communication with the fluid reservoir for evacuating fluid from the fluid reservoir via a fluid return channel separate from the fluid delivery channel.
The at least one fluid delivery port is at least one slit or channel in the material of the electrode carrier, or a semi-porous membrane that connects the fluid reservoir to the outer surface of the electrode carrier. The fluid reservoir and the at least one fluid delivery port may cooperate to preferentially deliver more therapeutic fluid closer to an apical end of the electrode carrier.
The placement of the micro-valve 101 typically, but not necessarily, requires drilling a hole approximately 0.8 to 2 mm or more in diameter on the promontory bone. The micro-valve 101 may be self closing, as shown in
When energy is delivered to the pump 111 to move fluid from the reservoir 113 down the catheter to the inner ear or when the fluid is moved via a spring loaded reservoir 401, the pressure is sufficient to open the micro-valve 101. When no energy or pressure is sensed by the micro-valve 101, the micro-valve 101 may close automatically, thereby sealing the inner from the middle ear. The micro-valve 101 closure may take place through the use of a titanium sphere 1001 attached to a spring 1003 on the inner ear side of the valve as shown in
In another with another embodiment of the invention, the micro-valve 101 may be securely connected directly to a screw-on canister 601 as shown in
As noted above, fluid delivery systems in accordance with the invention may be used in combination with an electronic prosthesis or implant, for example, a cochlear implant. This may be accomplished in two ways: fusion of the catheter and an electrode associated with the prosthesis or implant, or parallel delivery of fluid and electrical current to the body.
In accordance with various embodiments of the drug delivery system the reservoir or canister may be refillable.
As described above, the micro-septum connector 1203 is in fluid communication with an implantable fluid pump, a fluid port or reservoir, or an osmotic pump via a spear catheter 1205 and in fluid communication with the body of the subject via a port catheter 1207 which may be connected to or in fluid communication with another catheter or an electrode or electronic prosthesis (such as 1401). The fluid delivery device (such as the port catheter 1207 and electrode or electronic prostheses 1401) and the device that drives and delivers the fluid (such as the fluid pump 1201 or fluid port 1301) are designed to be implanted in a human subject or an animal subject in the course of a surgical procedure. The connection between the two devices is accomplished with the micro-septum connector 1203.
In one embodiment of the invention, the proximal end of the port connector 1501 and distal end of the spear connector 1503 do not join surface to surface. This is to prevent the creation of dead space between the flat surfaces of the micro-septum connector 1203 when joining the port connector 1501 and spear connector 1503 via the needle 1507. In such an embodiment, the needle 1507 of the spear connector 1503 traverses the septum 1505, but a remaining part of the needle 1507, anterior to the septum 1505, is exposed to body fluid and body tissue. Such a situation promotes a good tissue seal at the point where the needle 1507 enters the septum 1505. In addition, the encapsulating tissue is irrigated by the surrounding tissue and can respond well to any inflammation. It is also feasible to introduce tissue, fascia or muscle through the needle 1507 up to the flat end of the spear connector 1503. Introduction of tissue will promote good healthy tissue growth between the flat ends of the port connector 1501 and spear connector 1503. As can be seen in
An important feature of the port connector 1501 is the septum 1505. The septum 1505 is made preferably of rubber silicone. The port connector 1501 may also include a compression ring 1511. The compression ring 1511 (or other compression device) compresses the silicone to impart septum properties to the device. The compression ring 1511 is preferably made of medical grade titanium, however, any other material that can compress the silicone in a cylindrical part may be used. Such materials may include shape memory nitinol metal and memory shape polymer. The compression ring 1511 may be terminated toward the connecting side of the port connector 1501 by a guide or a guide mechanism 1513, stopper or other stopping device 1515 and locking mechanism 1517. A bacterial filter 1509 may or may not be placed between the port connector 1501 and the port catheter 1207. The port connector 1501 may also include a reservoir 1521 which may be lined with titanium shell or a titanium shell to prevent piercing by the needle 1507. The proximal end of the port catheter 1207 may optionally be silicone bonded with the port connector. A layer of silicone may be deposited on the entire port connector 1501 to prevent exposure of metal to the environment, and favor encapsulation. Deposition may be accomplished by dip coating the port connector 1501 in the appropriate silicone rubber solution.
The spear connector 1503 may be made of silicone or epoxy or any other bio compatible material as deemed necessary or profitable to the invention. A medical needle (such as 1507) of appropriate size, material, and shape is inserted in a mold such that both ends of the needle protrude out of the mold. Injection molding of silicone and/or medical grade epoxy solidly encases the core of the needle. A catheter (such as the spear catheter 1205) is introduced on one end of the needle and silicone is added and cured to seal the spear connector. The needle 1507 may be slanted and sharp on the side that will pierce the septum 1505. Note that the hole on the needle 1507 that will transfer fluid may be at the end of the needle or may be on the side of the needle at a short distance from the tip. As noted above, the micro-septum connector 1203 may also include a guide or guiding mechanism (such as 1513) to line up the needle 1507 and the port connector 1501 before piercing of the septum 1505 by the needle 1507. The guide or guiding mechanism 1513 permits the lining up of the tip of the needle 1507 with the center of the septum 1505. The guide or guiding mechanism 1513 also prevents large deviation of the needle 1507 upward or downward or sideways. Such deviation could lodge the tip of the needle 1507 in the internal wall of the port connector 1501 and prevent fluid flow.
The stopper or stopping device 1515 may be used to prevent the fusion of the spear connector 1503 and the port connector 1501 on their flat surfaces. The stopper or stopping device 1515 permits a section of the needle 1507 to be exposed to the body fluid even at full insertion. The stopper or stopping device 1515 consequently prevents the creation of a dead space between the flat surfaces of the implantable port and spear connector ends when connected (unless such a flat connection is profitable to the invention by, for example, having an antibiotic coating to prevent formation of a nidus of infection at all times).
The locking mechanism 1517 may be included in the micro-septum connector 1203 to promote the stability of the micro-septum connector 1203 under normal body movement and usage stress. The locking mechanism 1517 may be reversible to allow for replacement of one or more of the parts described above.
Once fabricated and sterilized by appropriate means, a surgeon may join the port connector 1501 and the spear connector 1503 by introducing the needle 1507 through the septum 1505 and, optionally, locking the micro-septum connector 1203 with the locking mechanism 1517. Before joining the port connector 1501 to the spear connector 1503, each may be filled separately with a fluidic pharmaceutical agent. Separate filling allows good priming of the implantable connectors 1501 and 1503 before connection. Filling of the spear catheter 1205 may be accomplished by filling the pump (usually though a pump septum), port with septum and reservoir, or osmotic pump. It may also be that the implantable connectors 1501 and 1503 are connected before any filling and priming of the devices takes place.
If removal and replacement of one or both of the implantable connectors 1501 and 1503 is desired, surgical intervention may include careful removal of tissue growth and membrane encapsulation around the port connector 1501 and removal of the spear connector 1503 by pulling back on one or the other connectors. At this stage either or both of the port connector 1501 and spear connector 1503 may be positioned in the biological environment of interest. This may be done after priming the system in the usual fashion described above. Once replaced and positioned, connection of the port and spear to connectors is accomplished by engaging the port connector 1501 and spear connector 1503 (perhaps by employing the guide 1513), piercing the septum 1505, and, if desired, locking the mechanisms via the locking mechanism 1517.
The system described with respect to
In accordance with the embodiment of
The fluid delivery systems of
The catheter 1801 may be conical or cylindrical in shape, round or elliptical in cross section and may have a rounded tip 1811 for ease of implantation. The catheter 1801 may be polymer based, and the polymer may include silicone, for good flexibility. Alternatively, the catheter may be made of a biodegradable polymer. Similarly, the catheter 1801 may be made of a material which shrinks when stretched.
The catheter 1801 may optionally include one or more reinforcing wires and/or ribbons 1807 made of hard polymer filament or metal or metal alloy to increase the pushability of the device and enhance implantation. The catheter 1801 may also include markers 1805 on the surface of the polymer to indicate insertion depth and/or an adjustable blocker to close the cochleostomy through which it is inserted. Embodiments of the catheter 1801 include double outlets 1809 to provide free flow in the fluid, and these outlets 1809 may be in opposite directions. Similarly, the catheter 1801 may include more than one channel 1803, 1817, 1819 in the center (or toward edges) of the catheter body to control a fluid or drug concentration-distribution profile. The catheter 1801 may also include a lubricating coating to enhance insertion. Similarly, the catheter 1801 may be coated with cortico steroid and/or antibiotics to prevent infection.
In related embodiments, the catheter 1801 may have an on/off switch or valve 1813 accessible by the subject (activated by a magnet or by mechanical pressure) located on the subject's body such as on the skin or, when used in connection with fluid delivery to the inner ear, on the skull between the fluid delivery reservoir and intra inner ear section. The valve or switch 1813 may be used to prevent backflow of fluid. The catheter 1801 may additionally include a moveable stopper 1815 to promote ease and accuracy of insertion. The catheter 1801 is designed with an internal channel 1803 for fluid delivery. For example, localized delivery of fluid to the inner ear may maintain spiral ganglions cell functional characteristics, regenerate dendrites, and promote the preservation of residual hearing, arrest progressive hearing loss. Applications may include delivery of cortico-steroids to prevent inflammation, medicine to arrest sclerosis, and tissue growth and be used for the novel treatment for tinnitus and vertigo.
Fluid delivery is accomplished through the hollow channel 1803 formed on the catheter lead up to a location intra scala. One or more outlets 1809 may be included in the catheter 1801. The channel 1803 may be connected to an internal micro pump or to a poll including a septum for external pumping of pharmacological agent. The hollow channel 1803 disposed close to the center or more off-centered to the edges of the catheter 1801 is formed by reverse molding. This means that a place holder may be included in a mold prior to injection molding. After injection molding, the place holder then is removed and a hollow channel is left in its place. Outlet(s) 1809 for the fluid delivery channel may be located basally and/or apically. The outlet(s) 1809 for fluid delivery may be coated with a ring of slow release bioactive agent to prevent tissue growth and occlusion of the outlets over time.
Each single outlet 1809 for fluid delivery may include two outlet channels 1817 and 1819 180 degrees apart. The two outlet channels 1817 and 1819 are connected either in a rectilinear fashion or they are offset from one another. The object of having the two outlet channels 180 degrees apart in a catheter designed for fluid delivery to the inner ear (as shown in
The electrode 2000 is designed to have properties which reduce the amount of force necessary for introduction in the cochlea. Reducing the electrode insertion force and increasing the electrode flexibility reduces the amount of trauma inflicted to the soft tissue which lines the scala tympani walls. Reducing the insertion trauma to the maximum is most beneficial to the patients who suffer from severe deafness and may be using a hearing aid in the ipsilateral ear, or have residual hearing which allows perception of low frequency sound unaided but have poor speech discrimination. The interest in keeping electrode trauma minimized is compounded by the fact that a patient implanted today may receive a device replacement or a device addition which restores some aspects of the degenerated neural pathway. If such neural a pathway is mechanically disturbed during electrode insertion there is a high likelihood that the pathway will be permanently destroyed.
A cochlear implant electrode is usually inserted through the inner ear (scala tymapani or scala vestibuli) through a hole drilled on the bony surface protecting the spiraling cochlea. If residual hearing is present it may be of interest to limit the insertion depth to a region below where acoustic hearing is present. A stopper 2015 on the electrode 2000 can limit insertion depth to a fixed predetermined value, 20 mm for example (but not limited to 20 mm) 20 mm corresponds to about 1 turn of cochlea. The stopper 2015 is designed to have vertical wall which prevents insertion beyond the cochleostomy. Slots may be built on the stopper 2015 to allow a surgeon to view the cochleostomy as the stopper 2015 approaches the external bone of the inner ear. In another embodiment, the stopper 2015 has a conical shape which allows for plugging of the cochleostomy. The stopper 2015 can also be a slider which is moved down from a superior region on the electrode 2000.
The insertion depth of the electrode 2000 may be controlled and limited to a pre-determined value. The pre-determined insertion depth value may be based on the patient audiogram. If the audiogram indicates significant residual hearing (50 dB or more for example) up to 2000 Hz, the surgeon could choose to limit insertion depth to 16 mm. The limitation of the insertion depth may take place with the use of a pre-cut biocompatible and sterile tube inserted from the from end of the electrode down to the stopper 2015. A 4 mm long tubing of sufficient thickness in front of a 24 mm long electrode (length from electrode tip to stopper wall) would limit insertion depth to 20 mm.
What distinguishes the electrode 2000 from other designs is the presence of the front end 2007 and back end 2005 on the electrode array 2003. The front end 2007 is much thinner than the back end 2005. In one embodiment, the front end 2007 of the electrode 2000 covers ¼ to ½ of the electrode extent. The bulk mass of front end 2007 of the electrode 2000 may be about ½ the bulk mass of the back end. It is understood that in this design the electrode 2000 neither grows continuously, nor is of constant diameter or cross sectional shape along the electrode extent. Rather, the electrode 2000 includes a discontinuity in its cross sectional shape. The discontinuity defines the limit of front end 2007 and the beginning of the back end 2005 of the electrode array 2003. The front end 2007 is designed to have low insertion and low bending forces required to push the array around the coiling, upward spiral geometry of the scala tympani. The back end 2005 is designed to maximize the pushability of the electrode to achieve a deep insertion when required. Pushability is important for electrode design since an electrode with low pushability will collapse around the cochleostomy before able to impart a forward movement to the tip 2011 of the electrode. To favor the insertion of the electrode 2000, the tip 2011 of the device may be thin and rounded with no sharp edges. In addition the front end 2007 and the back end 2005 of the electrode array 2003 may be tapered. Tapered in this sense means that the cross sectional area of the front end 2007 and back end 2005 grows continuously.
On the electrode extent, eight or more contacts 2009 are embedded or deposited on the polymer substrate. At present, eight contacts 2009 are the minimum required to reach asymptotic performance in speech understanding for implanted patients. The contacts 2009 may be made of platinum (Pt), platinum indium (PtIr), or indium oxide. The contacts 2009 may be round or oval or may be rectangular shaped with rounded edges. Rounded edges reduce the current density at the edges of the electrode contact. Current density at the edge of the contact 2009 surface is usually responsible for the initial contact dissolution of the metallic surface. The contacts 2009 may be in the form of a spherical ball such as that produced by flaming the tip of a platinum iridium wire. Each of the contacts 2009 may be a single or paired contact. In one embodiment, a combination of paired and single contacts may be used. Contacts 2009 located on the back end 2005 are paired while contacts located on the front end 2007 are unpaired. In this manner the flexibility of the front end 2007 of the electrode is preserved while the pushability of the back end 2005 is maintained.
Each contact 2009 is electrically connected to an insulated wire (2201, 2203, 2205, or 2207, shown in
Fluid delivery is applied through a hollow channel 2301 formed on the electrode lead up to a location intra scala. One to several outlets 2307 may be included between or close to the electrode contacts 2309. The hollow channel 2301 may be connected to an internal micro pump or to a poll including a septum for external pumping of pharmacological agent. The micro pump or the port may be located near the implant housing. The method for fabricating the hollow channel 2301 such that it is close to the center or more excentered to the edges of the electrode 2300 includes reverse molding. Again, in reverse molding, to form an internal hollow channel 2301, a place holder is included in the mold prior to injection molding. After injection molding the place holder is removed and a hollow channel is left in its place.
One or more outlets 2307 for the fluid delivery channel 2301 may be located near or in between basal contacts 2309 located on the electrode array 2303. The outlet(s) 2307 for fluid delivery may be coated with a ring of bioactive agent to prevent tissue growth and occlusion of the outlets over time.
Modeling of intra cochlear stimulation and animal EABR data indicates that an electrode array positioned close to the inner wall of the scala tympani would be beneficial to the neuro stimulation of cochlea implants. Such electrodes are referred to as perimodiolar electrode. There is a consensus that a perimodiolar electrode would lower psycho-acoustic threshold, increase the dynamic range of stimulation, and reduce channel interaction. Channel interaction may be caused by the field overlap from individual electrodes. Further potential benefits expected from a perimodiolar array include reduced power consumption to drive the implant, reduced side effects for the patient, implementation of innovative stimulation scheme, and better place coding of frequency. A larger number of electrodes may be effectively used.
For clarity the two segments are referred to as the electrode branch 2713 and a restraining arm 2711. The two segments 2711 and 2713 are and remain connected during the whole insertion process. The preferred method of connecting the segments is via the pressure mating of a rail molded on the electrode branch 2713 with a slot molded on the restraining arm 2711. In a cochlear implant, segments 2711 and 2713 are latter disconnected for the positioning of a section of the electrode branch against or close to the modiolus. The cochlea from a human temporal bone with the electrode and restraining arm in position is shown on
The restraining arm 2711 may include in its mass and along its whole length a platinum (Pt) or a platinum iridium (PtIr) ribbon or wire 2715, annealed or not annealed, to increases or decrease the rigidity of the restraining arm. Such control of the rigidity of the restraining arm 2711 is important in a cochlear implant to maintain good insertion properties (flexibility) as well as sufficient rigidity for when a retro positioning technique is applied to the electrode branch 2713 to displace the electrode branch 2713 closer to the modiolus. If the restraining arm 2711 is too soft, it will buckle during the retro-positioning technique.
The shape of the ribbon 2715 may be that of a rectangle with a ratio of length to width of 2 to 1 (as shown in
When used with a cochlear implant, the electrode 2700 may be sequentially inserted, and then the electrode 2700 is positioned toward the inner wall. In a first phase, the electrode array 2703 with the two segments 2711 and 2713 connected is inserted along the outer wall of the scala tympani. In a second phase, the section of the back end 2705 of the electrode array 2703 corresponding preferably to the basal turn of the scala tympani is displaced to come close to or to connect with the inner wall of the same scala tympani. This section is now referred to as the perimodiolar section. The perimodiolar section corresponds preferably to the basal turn of the cochlear because this is where the majority of electro-excitable neural elements are situated. These neural elements (spiral ganglion cells) would benefit the most from more proximal electrode stimulation. The remaining intra cochlear section of the electrode branch 2713 is referred to as the deep insertion section. The deep insertion section is designed to be deeply inserted in the scala tympani but it is not positioned against the inner wall by any voluntary action.
Following the full insertion of the segmented electrode array 2703 into the scala tympani of the cochlea, the restraining arm 2711 is held stationary posterior to the cochleostomy (outside the cochlea) by the surgeon and with some micro-tool such as forceps or tweezers. The electrode branch 2713 is then unmated or disconnected from the restraining arm and retracted out of the scala. This slight pulling of the electrode array 2703 out of the cochlea effectively uncouples the electrode branch 2713 and the restraining arm 2711, except at the point where the two segments converge. It is important to note that at the convergence point the two segments are attached via a metallic rod or ribbon 2715 made of PtIr 80-20%, for example, such as that supplied by Medwire Sigmund Cohn Corp, Mount Vernon, N.Y. In one embodiment, the end of the wire or ribbon 2715 fits into a silicone hollow cavity on the electrode branch 2713. Key to the retro-positioning technique is the synergy between the less flexible ribbon 2715 and wire in the core of the restraining arm 2711 and the more flexible electrode branch 2713. An important element of the electrode 2700 is the segmented aspect of the electrode. Another substantial element of the design is the option to connect firmly the two segments 2711 and 2713 for ease of insertion. The firm and yet detachable connection may be established by several means. One means of segment connection is via a rail and a slot having matching dimensions. The electrode branch 2713 and restraining arm 2711 may be pressure mated during manufacturing. The mating of the silicone keeps the electrode and restraining arm connected during insertion.
Another means of connecting the two segments 2711 and 2713 is via an envelope design. If such a design is adopted, the envelope may be round or ellipsoid in shape. It is understood that the mating of the electrode is not restricted to the designs shown and that any mating which is profitable for the connection, insertion, disconnection, and positioning of the electrode is feasible. In accordance with yet another method, the two segments 2711 and 2713 may be connected with a hydrogel which dissolves in the fluid of the inner ear within a few minutes of insertion. The binding of two dissimilar silicones which are disconnectable may also connect the two segments.
The electrode segments 2711 and 2713 have a convergence point so that when the implant needs replacement, the two segments 2711 and 2713 of the electrode array 2703 should be easily disconnectable. In order to achieve disconnectability as well as restraining action during retro-positioning, the two segments 2711 and 2713 may be joined by a bare PtIr ribbon 2715 section, which comes out of the restraining arm 2711 and is lodged snugly or loosely in an oriented silicone cavity molded on the electrode branch 2713. In case of revision surgery, the two segments 2711 and 2713 can be dislocated at their point of convergence by pulling back on the restraining arm 2711 with sufficient force. The cavity may be parallel to the axis of the array or may be oriented in such a way as to provide resistance for retro-positioning. The ribbon or wire 2715, which is used as the spine of the restraining arm 2711, may be terminated as a ball to reduce sharp edges. The two segments 2711 and 2713 of the electrode array 2703 may be attached together outside the cochlea. Such attachment may be advantageous to prevent the movement of the electrode branch 2713 in relation to the restraining arm 2711. With respect to a cochlear implant, movement of the electrode branch 2713 post-operatively could lead to a release of the connection of the electrode branch 2713 with the modiolus. The in such an embodiment, the two segments 2711 and 2713 may be attached with a closable titanium clip 2717 seen more clearly in
There are several advantages of the electrodes described above. First, in a cochlear implant, a section of the electrode may be deeply inserted in the cochlear, up to the apex, with minimized forces because of the front end design of the electrode. Additionally, a section of the electrode preferably corresponding to the first turn of the cochlear may be displaced toward to and up to the inner wall of the inner ear cavity. The two segments 2711 and 2713 of the electrode are and remain attached during the insertion process (but are disconnected during the positioning process, post insertion, and by voluntary action). The connection to the modiolus is independent of morphology and special tools are not required for insertion and positioning.
The front end 2707 of the electrode 2700, for up to a 15 mm length, may be coated with a thin biocompatible lubricating coating 2719. The coating 2719 may be permanent or biodegradable. Lubricating coating reduces the friction between the electrode and the tissue during insertion, therefore reducing insertion forces. Lubricating coating needs to be applied in a restricted front end length of the electrode so that instrument can hold the electrode and push in.
The electrode may also be equipped with a stopper (as show in
In yet another embodiment, the implantable electrodes may have an impermeable connector (as at 2313 of
As noted above, the reservoir 3103 may be metallic and silicone covered. The reservoir may also be conically shaped and the septum 3101 may be disposed on the greater diameter of the cone. When connected to the inner ear of a subject, such a conically shaped reservoir should be of dimensions adequate to snugly fit in a mastoidectomy, such that the non-septum side of the reservoir 3103 is connected to a catheter 3107. The catheter 3107 terminates on the outer side of needle 3109. As shown in
Embodiments of the present invention also include a cochlear electrode array based on the incorporation of a given amount of medicine into a portion of the silicone polymer elastomer that makes up the electrode body. Over time, the medicine is released from the elastomeric material and diffused into the fluid of the inner ear. The diffused molecules then target receptors of interest.
The inner ear presents various considerations for localized delivery of pharmacological agents which include drug targeting receptors suitable for the hearing organ such as neural tissue and soft tissue. The inner ear is a very small and essentially closed space so that any medicine released within the inner ear tends to remain confined within that space. Thus, any pharmacological agent which is slowly released in this environment tends to be bioactive only in the inner ear and there is very little diffusion outside of the inner ear.
The rate at which the pharmacological agent is released from the polymer matrix material of the electrode array depends on various factors. These include the amount of surface area of the drug eluting material which is exposed to the fluid surrounding the polymer. The drug releasing profile over time is directly proportional to the surface area of drug loaded polymer exposed to the surrounding fluids of the inner ear. The concentration of medicine within the polymer affects the length of the delivery time. The release rate of the pharmacological agent may also depend on other factors such as the cross-linked density of the material in the drug eluting portion and the volume of the drug eluting portion.
In the embodiments shown in FIGS. 36B-D, the electrode array includes a slot into which drug eluting material is incorporated. In
Examples of specific pharmacological agents suitable for post-surgical release into the inner ear include without limitation neurotrophic factors, gene therapy, anti-apotosis medicines, and anti-oxidants. Some medicines have neuro-protective effects and could help sustain the neural status of the inner ear after the somewhat traumatic cochlear implantation.
Other suitable pharmacological agents include anti inflammatory agents. For example, the saturated solubility in normal saline of the anti inflammatory agent may be not less than 26.4 μg/ml at 37° C. The electrode array may be adapted to release between 5 mg and 250 mg of anti inflammatory agent during the first week after implantation. The pharmaceutical agent may also include a bactericide.
Of special and immediate interest is the use of cortico steroids to control post-implantation fibrotic development. One example of such a cortico steroid is dexamethasone. For example, the electrode array may be adapted to release between 5 mg and 600 mg of dexamethasone during an initial 24 hour period of use. Other examples of cortico steroids suitable for use in a drug eluting cochlear implant electrode array include without limitation betamethasone, clobethasole, diflorasone, fluocinonide, triamcilione, or salt or combination thereof.
A drug eluting silicone material can be produced by first micronizing the pharmaceutical agent particles to a desired size. For example, the pharmaceutical agent may be in the form of solid particles of less than 200 μm mixed into the material of the drug eluting portion. The release rate of the pharmaceutical agent may be based on having particles of the pharmaceutical agent in a plurality of defined sizes. For example, in some embodiments, at least 90% of the particles may be less than 200 μm in size. In addition or alternatively, at least 50% of the particles may be less than 50 μm in size. The particles can be thoroughly mixed with liquid silicone polymer using a high speed homogenizer. In some embodiments, a cross-linking solution may be added to the mixture. The resulting mixture is injected into the space reserved for the drug eluting portion using a properly designed mold.
Concentration of the pharmaceutical agent in the surrounding inner ear fluid depends on the concentration and permeability of the pharmaceutical agent in the drug eluting material. The release time may be days to months depending on the cross linking density of the silicone, amount of loading of drug as a percentage of electrode array, volume of drug-loaded polymer, and surface area exposed to the fluid of the cochlea.
An electrode array according to an embodiment of the invention can be assembled in various steps. For example, the wires and electrode contacts used for electrical stimulation can be placed in one half of an array mold. A first stage of molding then encapsulates the wires and electrode contacts using a reverse molding or masking to leave a space where the drug eluting silicone material can be injected in a second step. This approach allows bonding of the two similar polymers to ensure a uniform contour of the electrode.
One advantage of using a two-stage molding process is that only a portion of the electrode array in the fluid of the inner ear need be loaded with a pharmaceutical agent. The extra cochlea portion of the electrode array can be made of non drug eluting material and need not participate in the drug release.
A multi-stage molding process involving multiple masking can also be used to successively add complimentary drug eluting material in more than one place, with each drug eluting portion having a different composition of pharmaceutical agent. In this manner, complimentary drugs or drugs targeting different receptors and at a different rate of diffusion can be incorporated in the electrode array.
Polymer rods loaded with a pharmacologic agent may be prefabricated. The rod of drug eluting material may be made of a silicone of the same or similar composition as that used in the fabrication of the main non drug eluting portion of the electrode array. For example, drug eluting rods can be prefabricated in a high level pharmaceutical lab equipped with the necessary instrumentation. The rods can then be shipped to be assembled with the cochlear implant electrode array at another location. For example, the electrode arrays shown in
As noted above, the invention and its embodiments described herein are not limited to application to the inner ear. Other applications, such as anywhere in the body where it is desirable to have a pump and a delivery catheter with or without electrical stimulation, are also possible with the use of this connecting system. For instance, it may be that such a connection is made on or in the skull at a preferred location for fluid delivery to some location in the brain.
Embodiments of the present invention are also directed to an implantable electrode carrier that can accommodate a variety of therapeutic fluids such as a therapeutic drug in liquid form. The drug delivery function is integrated into the electrode carrier without changing the size or shape of the carrier, and without increased risk to the patient. The therapeutic fluid can be dispensed from the carrier without changing the pressure or volume of the surrounding tissue, for example, to cochlear fluid within the cochlea. Although illustrative embodiments are described with respect to the cochlea and cochlear implants, the invention is not limited such a specific application and may be useful in other implantable systems at other locations within the body.
In another embodiment, the delivery ports may be one or more slits or channels in the material of the electrode carrier 3901, which, for example, may be carved by lasering such as with a femto laser. The release profile for the therapeutic fluid in the fluid reservoir 3905 to the tissue surrounding the electrode carrier 3901 is determined by the dimensions—length, width, thickness—of the delivery ports 3906. The release time (days, weeks, etc.) is determined by the size of the fluid reservoir 3905. Specific embodiments may be used as a passive drug delivery system using diffusion of reservoir molecules, or as an active delivery system using a pump to deliver the therapeutic fluid from the fluid reservoir 3905 out the delivery ports 3906.
In one embodiment, the fluid reservoir 3905 may be filed with elastomeric material such as silicone which includes a dispersion of therapeutic molecules. These molecules may be released over time from the elastomeric material in the fluid reservoir 3905, out through the delivery ports 3906 to the surrounding tissue.
While the invention has been described in connection with specific embodiments thereof, it will be understood that it is capable of further modification. This application is intended to cover any variation, uses, or adaptations of the invention and including such departures from the present disclosure as come within known or customary practice in the art to which invention pertains.