CA2616878A1

CA2616878A1 - Cardiac harness having an optimal impedance range

Info

Publication number: CA2616878A1
Application number: CA002616878A
Authority: CA
Inventors: Alan Schaer; Matthew G. Fischler
Original assignee: Paracor Medical, Inc.; Alan Schaer; Matthew G. Fischler
Current assignee: Paracor Medical Inc
Priority date: 2005-08-01
Filing date: 2006-07-14
Publication date: 2007-02-08
Also published as: EP1922021A1; JP2009502418A; WO2007015762A1; US20070027516A1; US7587247B2

Abstract

A system for treating the heart including a cardiac harness (10) configured to conform generally to at least a portion of a patient's heart. The system also includes an electrode (18) associated with the cardiac harness and positioned on or proximate to the epicardial surface of the heart. In order to ensure that the electrode will operate with a pulse generator, the system has an impedance between approximately 10 ohms and approximately 120 ohms.

Description

CARDIAC HARNESS HAVING AN OPTIMAL IMPEDANCE RANGE
Background of the Invention 1. Field of the Invention:
The present invention relates to a device for treating heart failure. More specifically, the invention relates to a cardiac harness having electrodes for providing defibrillation and/or pacing/sensing therapies. The design of the cardiac harness provides electrodes integrated with the cardiac harnless having an impedance that optimize the coinpatibility of the system with cominercially available internal cardioverter defibrillators.
2. General Background and State of the Art:
Cardiac harnesses, such as those disclosed in U.S. Serial No. 10/704,376 ("the '376 application"), may be used to treat cardiac heart faih.ire. The entire contents of the '376 application is incorporated herein by reference. To treat other heart failures, including cardiac arrhythinias, the cardiac harness of the '376 application may include electrodes that are con.nected to an iinplantable cardioverter defibrillator ("ICD"), which are well known in the art. Such electrodes are capable of delivering a defibrillating electrical shock from the ICD to the heart. These electrodes may also provide pacing/sensing fiuzctions to the heart to treat cardiac failures, including bradycardia and tachycardia.
It is desirable to have the cardiac harness with electrodes be coinpatible with coininercially available ICDs and defibrillation capable cardiac resynchronization therapy ("CRT-D") and pulse generators ("PG"), such as those fiom Guidant, Medtronic, and St.
Jude Medical. In order to be coinpatible with these commercially available ICDs and CRT-D PGs the electrodes of the cardiac harness inust have an appropriate electrical iinpedance. If the system (cardiac hanless with electrodes connected to a power source) has an impedance that is too low, the systein could become damaged. On the other hand, if the system has an iinpedance that is too high, the system may produce an insufficient ainount of electric current to travel across the cardiac tissue to sufficiently depolarize a critical amount of cardiac tissue to result in termination of the fibrillating wavefionts.
Therefore, what is needed is a cardiac harness having defibrillation and/or pacing/sensing capabilities, wherein the electrodes of the cardiac harness have an impedance that is witllin.
an appropriate range.

Stuninary Of The Invention In accordance wit11 the present invention, a system for treating the heart includes a cardiac harness configured to conform generally to at least a portion of a patient's heart.
The systein also includes at least one electrode associated with the cardiac harness and positioned proximate to an outer surface of the heart, and a power source in coinmunication with the electrode. The electrode and power source are at least a part of an electrical circuit. The electrical circuit may also include a condtictor in coinmunication between the electrode and the power source or the electrode and power source may coininunicate wirelessly. Iii order to ensure that the electrical circuit will function properly, the electrical circuit has an impedance between approximately 10 ohins and approximately 120 ohins. It is even more preferred that the impedance range be between approximately 20 oluns and 80 ohms. The lower impedance range is dictated by the fitnctionality of the power source or pulse generator. Having too low of an impedance (under 10 oluns) can damage the electrical circuit incorporated with the cardiac hanless.
The upper iinpedance limit is that which continues to provide an adequate defibrillation threshold ("DFT").
Several alterations can be made to the system to increase its inipedance and avoid falling tuider the lower impedance limit of 10 oluns. In one aspect, a dielectric material such as silicone rubber is disposed on a pericardial side of the electrode (side of electrode facing away from the heart), leaving an epicardial side of the electrode (side of electrode in contact with the heart) un-insulated. Insulating the pericardial side of the electrode increases the impedance of the system, and prevents the system from having an impedance that falls under the lower iinpedance limit.
lii another aspect, the pitch of a normal electrode coil can be increased.
Increasing the pitch of the electrode coil decreases its surface area, and consequently, increases the impedance of the system.
In yet another aspect of the present invention, the coinposition of the conductive wire or conductor, which may include an MP35N-Ag coinposite, can be altered by changing the silver content. The preferred balance of iinpedance and mechanical strength is achieved with a 25% silver content of the conductive wire composite. In order to keep the impedance of the present system above the lower impedance limit, the silver content within the conductor can be froni 0% to about 50%.

Also, the cross-section of the wire forming the electrode can be reduced to increase the impedance. In this embodiment, changing the wire of the electrode in any way to reduce the area of its cross-section or its outer diameter will increase its impedance. The width and/or heigllt of the cross-section of the wire forming the electrode can be reduced to decrease its cross sectional area. In anotller einbodunent, the cross-sectional shape of the electrode coil wire may be changed to reduce its surface area. In one instance, the wire of the electrode can be changed from a rectangular cross-section to a circular cross-section.
Further, the overall outer diameter of the electrode can be reduced to increase the iinpedance of the system. If the electrode is in the form of a helical coil, the wire fonning the coil can be wound tighter to decrease the overall outer diaineter of the helical coil.
In a fitrtlier aspect, a resistor can be plugged in-line with the lead system to increase the impedance of the system.
A.iother aspect includes an electrode with circumferentially insulating seginents disposed along its length. The insulating segments can be fonned of any dielectric material such as silicone rubber, and may be any size. Further, any number of insulating segments may be disposed along the electrode. The insulating seginents disposed around the electrode reduce the exposed surface area of the electrode, thereby increasing the impedance. The insulating segments may also force a redistribution of current in the exposed regions of the electrode in order to optimize the DFT.
Anotller aspect includes an electrode with a resistive film (i.e., an oxide layer) disposed on the electrode surface. The resistive film could fiirther be deposited non-unifonnly so as to spatially modulate surface resistance (i.e., to reduce current density edge effects, or to alter the current distribution along the length of the electrode to optimize the DFT).
In yet anotlier aspect, the lengtll of the electrode can be shortened. By shortening the electrode, the overall surface area of the electrode is decreased, thereby increasing the iinpedance of the system.
Brief Description of the Drawings FIG. 1 is a perspective view of a cardiac haniess including a lead system that is coiulected to a power source.
FIG. 2 is a cross-sectioarial view talcen along line 2-2 of FIG. 1.
FIG. 3 is a partial cross-sectional view of a distal end of an electrode attached to a cardiac hanless.

FIG. 4 is a cross-sectional view talcen along line 4-4 of FIG. 1 showing an electrode having its pericardial side insulated.
FIG. 5 is a partial view of a helical coil of an electrode with the pitch of the winding increased.
FIG. 6a is a cross-sectional view of a wire forming an electrode with reduced dimensions.
FIG. 6b is a cross-sectional view of a wire forming an electrode with less cross-sectional area due to the change in the cross-sectional shape of the wire.
FIG. 7 is a partial view of a resistor plugged in-line witll a conductor wire.
FIG. 8 is a partial view of an electrode having circumferential seginents of a dielectric material disposed along the lengtli of the electrode.
Detailed Description of the Preferred Embodiments The present invention is directed to a cardiac harness system for treating the heart.
The cardiac hanzess system of the present invention couples a cardiac harness for treating the heart witl-i a cardiac rllytlun management device. More particularly, the cardiac harness includes rows or undulating strands of spring elements that provide a coinpressive force on the heart during diastole and systole in order to relieve wall stress pressure on the heart. Associated with the cardiac harness is a cardiac rhytlun management device for treating any number of irregularities in heart beat due to, ainong other reasons, congestive heart failure. Tlzus, the cardiac rhytlun management device associated with the cardiac hanless can include one or more of the following: an iinplantable cardioverter defibrillator ("ICD") with associated leads and electrodes; a cardiac paceinalcer (or cardiac resynchronization therapy ("CRT") pulse generator) inchiding leads and electrodes used for sensing cardiac function and providing pacing stimuli to treat synchrony of bot11 vessels; and a combined ICD and pacemaker (referred to as a ("CRT-D")), with associated leads and electrodes to provide a defibrillation shock and/or pacing/sensing functions.
The cardiac hanless system may include various configurations of panels connected together to at least partially suinound the heart and assist the heart during diastole and systole. The cardiac harness system also includes one or more leads having electrodes associated with the cardiac haniess and a source of electrical energy supplied to the electrodes for delivering a defibrillating shock or pacing stimuli.
In one en7bodiment of the invention, as shown in FIG. 1, a cardiac hanless 10 includes four panels 12 of generally continuous undulating strands 14. A panel includes rows or undulating strands of hinges or spring elements that are connected together and that are positioned between a pair of electrodes, the rows or undulations being highly elastic in the circumferential direction and, to a lesser extent, in the longitudinal direction.
The cardiac hainess also includes separate leads 16 having conductive electrode portions 18 that are spaced apart and which separate the panels 12. As shown in FIG. 1, the electrodes are fonned of a conductive coil wire, preferably in a helical maiuier. A
conductive wire or conductor 20 is attached to the coil wire and to a power source 22, fonning a part of the electrical circuitry of the system. As used herein, the power source can include any of the following, depending upon the particular application of the electrode: a pulse generator ("PG"); an ICD; a pacemaker or CRT; and an iinplantable cardioverter defibrillator coupled with a pacemaker or CRT-D. In the einbodiment shown in FIG. 1, the electrodes are configured to deliver an electrical shock, via the conductive wire and the power source, to the epicardial surface of the heart so that the electrical shock passes through the myocardium. The electrodes can be spaced so that they are about 0 apart, 45 apart, 60 apart, 90 apart, 120 apart, or any arbitrary arc length spacing, or, for that matter, essentially any arc length apart around the circuinference of the heart in order to deliver an appropriate electrical shock. As previously described, it may become necessary to defibrillate the heart and the electrodes 18 are configured to deliver an appropriate electrical shock to defibrillate the heart.
Iii the embodiment shown in FIG. 1, a Y-junction member 21 is used to join two adjacent conductor wires 20. As best shown in FIG. 2, the Y-junction is a low-profile molding of silicone rubber or other dielectric material having two lumens 23, one for each conductor wire. Any number of lumens may be formed within the Y junction to join more than 2 conductor wires. In this einbodiment, the proximal ends of the joined conductors are crimped together into a pin (not shown) that is attached to the power source 22. The molding that fonns the Y junction meinber may extend from the Y junction to the power source, or may only extend a certain distance that ends before the power source. The Y-junctions helps to organize and nlanage the conductors witliin a patient's body. In otlier embodiments, the conductors may not be joined together with the Y junction niember.
As best shown in FIG. 3, the electrodes 18 are attached to the cardiac harness 10, and more particularly to the undulating strands 14, by a dielectric material 24. The dielectric material insulates the electrodes fiom the cardiac harness so that electrical cturent does not pass from the electrode to the hai7iess thereby undesirably shunting current away from the heart for defibrillation. Preferably, the dielectric material covers the undulating strands and covers at least a portion of the electrodes 18. FIG. 3 also shows in inore detail how the conductive wire or conductor 20 is in communication witli the electrode 18. In the embodiment shown, the electrode portion is a ribbon of conductive material that is coiled around and welded to a dome 26 at a distal end of the cardiac haniess. The dome is also formed of a conductive material (such as MP35N) and has a distal end with a blind hole 28, and a proximal end forming a seat 30. During manufacturing, silicone rubber or another dielectric material flows into the blind hole 28 to help attach the dielectric material at the end of the electrode. Also during assembly, a distal end of the conductor wire 20 is placed and crimped within the seat 30, thereby placing the conductor in electrical cormmnunication with the electrode 18 via the dome 26.
In this embodiment, the contact junction between the conductor and the electrode is at the distal end of the cardiac harness where there is less bending moments, and therefore, it is less likely that this contact junction will fracture or fatigue. FIG. 3 shows the dielectric material 24 molded around the ends of the undulating strands 14, and a cap 32 disposed at the end of the undulating strand. Grip pads (not shown) may also be attached to the dielectric material to help hold the cardiac harness in place once positioned around a potion of a beating heart.
The cardiac harness 10 may be produced in a range of sizes, with distinct lengths depending on the size and the number or rows of undulating strands 14. In the embodiment shown in FIG. 1, the cardiac harness includes six rows of undulating strands, however, other einbodiments may include fewer or more rows of undulating strands. The electrode 18 lengtll and surface area is preferably propoi-tional with the hanless length. For exainple, the length and surface area of the electrode can be approximately 49 inin and 307 mm2, 65 mm and 407 mm2, and 81 mm and 505 imnZ for a cardiac hanlesses having four, five, and six undulating strand rows, respectively. However, the size of the electrode may remain constant regardless of the size of the cardiac hanzess.
Iii one embodiment, the cardiac harness 10 is intended to fiuiction with conunercially available pace/sense leads and ICD pulse generators. To ensure the cardiac harness is compatible witli cominercially available ICD and CRT-D pulse generators, it inust have an appropriate electrical iinpedance. Coinmercially available ICD
and CRT-D
pulse generators, such as those from Guidant, Medtronic, and St. Jude Medical, typically have a lower impedance limit below which the device will not deliver a shock during progranuned device testing at implantation. This limit, typically 20 52, is dictated by the current carrying limits of the intemal pulse generator circuitry. Since the ICD delivers a set voltage from a charged capacitor, as the system iinpedance drops, the delivered current increases. Once implanted, the ICD should deliver a defibrillation shock even if the impedance drops below 20 Q, although there is a risk that the circuitry of the system will be dainaged. Depending on the initial voltage, actual unit damage may not occur until about 10 SZ. Therefore, it is preferred that the lower iinpedance range of the lead system attached to the cardiac harness is no lower than about 20 S2, with a fiinctional limit of about 10 Q.
Several paraineters affect the system impedance. These include, but are not limited to, the iiiherent resistivity of the tissue voluine through which the defibrillation current flows (may be affected by tissue density, tissue fluid levels, air volume, etc.); the distance between the electrodes attached to the cardiac harness; the surface area of the electrodes exposed to the body tissues; the electrode geometiy (and impact on current edge effects);
the inter-relationship between isopotential lines of current flow; the resistance in the lead electrodes, conductors, and contact junctions, and ICD or CRT-D circuitry;
electrode material (polarization effects) and microscopic surface texture (i.e., fiactal coatings, black Pt, etc.); and the morphology of the shock wavefonn (i.e., repolarization effects of a biphasic waveform).
As the length of the electrode 18 increases to extend along cardiac hai7lesses of varying lengths, the impedance of the system decreases. In other words, the larger cardiac hanless have longer electrodes with more exposed surface area than the electrodes attached to smaller cardiac harnesses, and the electrical circuitry associated witll the longer electrodes also have a lower iinpedance than the electrical circuitry associated with the smaller electrodes. Therefore, w11at is needed is a way to increase the impedance of the system to avoid falling under the lower impedance limit of 20 E2. In one einbodiment as shown in the cross-sectional view of FIG. 4, dielectric material such as silicone iLibber 34 is disposed on a pericardial side 36 (side of electrode facing away from the heart) of the electrode, leaving an epicardial side 38 (side of electrode in contact with the heart) of the electrode un-insulated. Any length of the pericardial side of the electrode may be insulated up to the entire length of the electrode. Insulating the pericardial side of the electrode increases the system iinpedance, and thereby prevents the systein from having an inzpedance that falls under the lower impedance limit. Althougli not preferred, it has also been contemplated that a certain portion of the epicardial side of the electrode could be insulated in addition to or instead of the pericardial side to reduce the electrodes surface area and increase its impedance.

In another embodiinent, the pitcli of electrode coil 18 can be increased. The coil shown in FIG. 5 has a greater pitch coinpared to the pitch of the electrode shown in FIG. 1.
hicreasing the pitch of the electrode coil decreases its total surface area per unit length, and consequently, increases the systein impedance.
In yet anotl-ier embodiment, the composition of the conductive wire or conductor 20, which may include an MP35N-Ag coinposite, can be altered by changing the silver content. By specifying tlie silver content of the conductor to be around 25%, a prefelTed balance of iinpedance and mechanical strengtll of the lead system is achieved.
In order to keep the iinpedance of the present system above the lower impedance limit, the silver content witllin the conductor can be from 0% to about 50%.
The cross-sectional dimensions of the wire forining the electrode coil 18 can be reduced to increase the impedance. In this embodiment, changing the wire of the electrode in any way to reduce the area of its cross-section or its outer diameter will increase impedance. The width and/or height of the wire fonning the electrode coil can be reduced to decrease its cross sectional area as shown in FIG. 6a, where the dotted line represents the electrode before the reduction. Also, in another einbodiinent as shown in FIG. 6b, the cross-sectional shape of the electrode coil wire may be changed to reduce its area. In this instance, the wire of the electrode was changed from a rectangular cross-section to a circular cross-section. In other embodiments, the cross-sectional shape may be changed to an any shape giving the electrode wire a lesser cross-sectional area, such as oval or any polygonal shape.
In other embodiments, the overall outer diameter of the electrode can be reduced to increase the impedance of the system. If the electrode is in the fonn of a helical coil, the wire fonning the coil can be wound tighter to decrease the overall outer diaineter of the helical coil, and thereby decreasing the overall surface area of the electrode.
In a fi.u-tlzer embodiment, a resistor 40 can be plugged in-line with the lead system to increase the iinpedance of the system. FIG. 7 is a partial view of one conductor 20, showing the resistor 40 plugged in-line with the conductor. A separate resister can be plugged in-line with each condtictor of the system. The conductor 20 is usually insulated with a dielectric material 24, and as shown in FIG. 7, it is preferred that resistor also be insulated with a dielectric material.
Referring now to FIG. 8, another embodiment is shown where the electrode 18 includes circumferentially insulating seginents 42 disposed along its length.
Only the electrode is shown in this figure for clarity reasons, with three separate insulating seginents 42 disposed coinpletely around the electrode. The insulating segments can be forined of any dielectric material such as silicone rubber, and may be any size, up to the length of the electrode. Furtlzer, any number of insulating seginents may be disposed around the electrode, including 1, 2, 3, 4, 5, etc., insulating segments. The insulating seginents can also be equally spaced apart from another, or in other embodiments, can be randomly spaced apart. The insulating seginents disposed around the electrode reduce the exposed surface area of the electrode, thereby increasing the impedance.
In another einbodiment, the electrode 18 may include a resistive film (i.e., an oxide layer) disposed on at least a portion of its surface. The resistive fihn could further be deposited non-unifonnly so as to spatially modulate surface resistance (i.e., to reduce current density edge effects, or to alter the current distribution along the length of the electrode to optimize the DFT). By disposing the resistive film along the surface of the electrode, the impedance of the system will increase.
hi yet another embodiment, the length of the electrode 18 can be shortened.
For example, the length of the electrode shown in FIG. 1 could be shortened to decrease the surface area of the electrode. By shortening the electrode, the overall surface area of the electrode is decreased, thereby increasing the impedance of the system.
The present system inust also not exceed an upper impedance level. If the impedance of the system is too high, an insufficient amount of current will travel across the cardiac tissue to sufficiently depolarize a critical ainount of cardiac tissue to result in tennination of the fibrillating wavefronts. Witli biphasic wavefonns, studies suggest that a voltage gradient of at least 3V/cm is required to achieve 80% defibrillation success. See Zhou X, Daubert JP, Wolf PD, Smith WM, Ideker RE; Epicardial Mapping Of Vetricular Defibrillation With Monophasic And Biphasic Shocks In Dogs; Circulation Research 72:145-160 (1993); which is hereby incorporated by reference. So, while there is no pai-ticular upper impedance limit, the iinpedance needs to be witllin a reasonable range to ensure defibrillation success. One way to define a reasonable upper limit is to first determine what impedance values are typical in commercially available devices that have acceptable DFT values.
The typical system shock impedance values seen in htunans have been reported in various studies (see table shown in Appendix 1). The data froni the table of Appendix 1 was gathered from the following references, also listed in Appendix 1; 1) Rinaldi A.C., Simon R.D., Geelen P., Reelc S., Baszko A., Kuehl M., Gill J.S., A
Reznelo7nized Prospective Study Of Single Coil Versus Dual Coil Defibrillation In Patien.ts With VentYicular Arrhythinias Undergoing Iinplantable CaNdioverter Defibrillator Thef=apy, Journal of Pacing and Clinical Electrophysiology 26:1684-1690 (2003); 2) Gold MR, Olsovsky MR, Pelini MA, Peters RW, Shorofslcy SR, Connpar=ison Of Single And Dual Coil Active Pectoral Defibyillation Lead Systenzs, Jouinal of the American College of Cardiology 1391-4 (1998); 3) Schulte B, Sperzel J, Carlsson J, Schwarz T, Ehrlich W, Pitscluzer HF, Neuzner J, Dual-Coil Vs. Single-Coil Active Pectoral Innplantable Defibrillator Lead Systefns: Defibrillation Energy RequiNen2ents And Probability Of Defibrillation Success At Multiples Of The Defibrillation Energy Requiyernen.ts, Europace 3:177-180 (2001); 4) Sandstedt B, Kemiergren C, Edvardsson N, Bidirectional Defibrillation Using Irnplantable Defibrillators: A Prospective Randonaized Compaerison Between Pectoral And Abdominal Active Generators, Journal of the American College of Cardiology 1343-1353 (2001); and 5) Shorofsky SR, Peters RW, Rashba EJ, Gold MR, Co771pai ison Of Step-Down And Binary SearchAlgoritl2ins For Deternain.ation Of Defibrillation. TI1.yeshold In Hunaans, Journal of Pacing and Clinical Electrophysiology 27:218-220 (2004). All of these references are herein incorporated by reference.
Based on the data from the above references, the mean iinpedance at iinplant for a dual coil active pectoral PG system is about 40 0 (standard deviation ranges 4-1052), and about 6052 10n for a single coil active PG system. The single (distal) coil used in these studies was about 50 mm long and had a surface area of about 450-480 mmZ. The second (proximal) coil in the dual-coil systeins was about 72 inm long and had a surface area of about 660-671 mm2.
To compare, a study in pigs was conducted to determine the DFT at the time of implantation of one embodiment of a cardiac harness having four rows of undulating strands and with 60 intra-electrode spacing. The electrodes incorporated with the cardiac harness used in this experiment had an exposed inner and outer coil surface with a surface area of about 660 mmz. The results from this study are presented in U.S.
Serial No.
11/051,823 ("the '823 application"), which is hereby incoiporated by reference in its entirety. In one experiment, the a defibrillation vector for the defibrillating cardiac hanless system was created from the right ventricular electrodes of the cardiac hailless to the left ventricular electrodes of the cardiac hanless and the active can coupled together. For this experiment, as listed in the '823 application, the mean DFT was 9.6 J and the inipedance was nleasured at 27 SZ. Also listed in the '823 application were comparable values for the mean DFT and impedance from a standard single lead defibrillation coil in the right ventricular endocardium, with a defibrillation vector from the defibrillation coil to the active can. The mean DFT was detennined to be 19.3 J and the impedance was measured at 46 0. Coinpared with the human data from a similar system reported in Appendix 1, the mean DFT values of the pig experiment with the defibrillation vector from the defibrillation coil disposed in the right ventricular endocardium to the active can are about 8 J higher and the impedance slightly lower. Also of note in the pig study was the advantage of increasing the intra-pair electrode spacing in lowering the mean DFT.
As wit11 other coinmercially available epicardial patches and, to some extent, endocardial leads, it is anticipated that the impedance of the iinplant will change with time after implantation. See Schwartzman D, Hull ML, Callans DJ, Gottlieb CD, Marchlinski FE; Serial Defibrillation Lead Iinpeelarace In. Patients Tvith Epicardial And Nontho7 acotomy Lead Systeyras; Journal of Cardiovascular Electrophysiology 7:697-703 (1996), which is hereby incoiporated by reference. Thus, wlien designing the cardiac haniess implant to function witll an ICD or CRT-D systein, consideration of the time course of iinpedance change is important to ensLUe the system remains fiulctional througllout the healing phase.
In order to test a cardiac harness having six-rows of undulating strands, additional bench-top tests were conducted in a saline taiilc with the cardiac harness including defibrillation electrodes placed over a saturated heart-shaped piece of foain (to mimic a human heart). Shock tests on a cardiac hanzess including defibrillation electrodes, which were exposed or un-insulated on both sides of the electrode, and having four-rows of undulating strands were performed. The defibrillation vector of this test simulated the vector from the right ventricular pair of electrodes to the left ventricular pair of electrodes coupled to the active can in the left pectoral region. During this test, the impedance was measured at about 26 S2 (similar to the pig data referenced above). Repeating the test with the six-row cardiac hanless including defibrillation electrodes with 60 intra-electrode spacing, and imzer and outer coil surface exposed giving an electrode surface area of about 1060 mmZ per pair, resulted in an impedance of about 20 52, which is less than the impedance of the smaller cardiac harness.
Because of the concenl that the six-row cardiac hanless including defibrillation electrodes would have an iinpedance too close to the lower limit of the ICD, the design of the cardiac haniess was altered by adding silicone ilibber insulation to the outside (pericardial side) of the electrodes, leaving only the inside surface (or epicardial side) exposed. This resulted in an exposed electrode surface area of the four-row and six-row pairs of 330 iYuna and 530 mm2, respectively. The expectation was that by reducing the electrode surface area, the impedance would increase. A repeat of the above in-vitro tests resulted in the four-row cardiac harness having its iinpedance increase from about 26 Q to about 39 Q, and the six-row cardiac hanless having its impedance increase from about 20 SZ to about 30 Q. A coinparison of 60 and 45 intra electrode separation showed no significant difference in the impedance level.
While insulating the outside of the electrode was one way to increase impedance, other methods, such as those discussed above can also be used to increase or otherwise inodify the systein shock impedance.
Again, the lower iinpedance range is dictated by the fiuzctionality of the power source or pulse generator. This is preferably no lower t11an about 20 52,, with a functional limit of about 10 S2. The upper impedance limit is that which continues to provide an adequate DFT. Given the data in humans discussed above, the prefelTed upper impedance range is about 80 Q. However, as noted in the pig study, the cardiac hanless with defibrillating electrode geometry may provide a more iulifonn distribution of ctuTent compared to coininercial leads, and tllerefore may be able to provide adequate voltage gradients with higher impedance values than are reported with conventional electrodes.
Thus, the fiinctional impedance range is estimated to rLU1 about 50% higher, up to 120 Q.
In sLuninary, the preferred impedance range for the cardiac harness with lead system is about 20 SZ to about 80 52,, with a functional range of about 10 Q to 120 Q.
Although the present invention has been described in tenns of certain prefeired embodiments, other einbodiments that are apparent to those of ordinary skill in the art are also within the scope of the invention. Accordingly, the scope of the invention is intended to be defined only by reference to the appended claims. While the impedance values, electrode dimensions, types of materials and coatings described herein are intended to define the paraineters of the invention, they are by no means limiting and are exemplary embodiments.

APPENDIX

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CD
o Appendix 1 1) Rinaldi AC, Simon RD, Geelen P, Reek S, Baszko A, Kuehl M, Gill JS, A
Randon2ized Prospective Study Of Single Coil Versus Dual Coil Defibrillation In Patients With.
Ven.tricular Arrhythmias Un.deNgoing Implantable Cardioverter Defibrillator TheYapy, Journal of Pacing and Clinical Electrophysiology 26: 1684-1690 (2003);
2) Gold MR, Olsovsky MR, Pelini MA, Peters RW, Shorofsky SR, Comparison Of Single And Dual Coil Active Pectoral Defibrillation Lead Systems, Journal Of The Ainerican College Of Cardiology: 1391-4 (1998);
3) Scllulte B, Sperzel J, Carlsson J, Schwarz T, Ehrlich W, Pitschner HF, Neuzner J, Dual-Coil Vs. Single-Coil Active Pectoral Implantable Defibrillator Lead Systeins:
Defibrillation Energy Requirements And Pr obability Of Defibf illation Success At Multiples Of The Defibrillation Energy Requif enaents, Europace 3:177-180 (2001);
4) Sandstedt B, Keimergren C, Edvardsson N, Bidirectional Defibrillation Using Implantable Defibrillators: A Prospective Randomized Comparison Between.
Pectoral And AbdominalActive Generators, Journal Of The Atnerican College Of Cardiology:
24:1343-1353 (2001); and 5) Shorofsky SR, Peters RW, Rashba EJ, Gold MR, Compar=ison Of Step-Down. And Binary Seay ch Algoritlams For Determination Of Defibi illation Tlareshold In Humans, Jotunal of Pacing and Clinical Electrophysiology 27:218-220 (2004).

Claims

1. A system for treating the heart, comprising:
a cardiac harness configured to conform generally to at least a portion of a patient's heart;
an electrode attached to the cardiac harness and positioned on or proximate to the epicardial surface of the heart;
a power source in communication with the electrode, the electrode and power source are at least a part of an electrical circuit; and the electrical circuit having an impedance between approximately 10 ohms and approximately 120 ohms.

2. The system of claim 1, wherein the electrical circuit having an impedance between approximately 20 ohms and approximately 80 ohms.

3. The system of claim 1, further comprising a conductor in communication with the electrode and the power source.

4. The system of claim 3, further coinprising a resistor disposed in-line with the conductor.

5. The system of claim 1, wherein the electrode includes an epicardial side opposite a pericardial side, at least a portion of the epicardial side of the electrode being insulated with a dielectric material.

6. The system of claim 1, wherein the electrode includes an epicardial side opposite a pericardial side, at least a portion of the pericardial side of the electrode being insulated with a dielectric material.

7. The system of claim 1, wherein the conductor includes less than about 50%
silver.

8. The system of claim 1, wherein the electrode includes at least one segment of a dielectric material disposed circumferentially around the electrode, and the at least one segment of dielectric material has a length shorter than the length of the electrode.

9. A method of increasing the impedance of electrodes associated with a cardiac harness, comprising:
decreasing surface area of an electrode attached to a cardiac harness configured to conform generally to at least a portion of a patient's heart.

10. The method of claim 9, wherein decreasing the surface area of the electrode includes coating the electrode with a dielectric material.

11. The method of claim 10, wherein coating the electrode with a dielectric material on a pericardial side of the electrode.

12. The method of claim 10, wherein coating the electrode with a dielectric material on an epicardial side of the electrode.

13. The method of claim 10, wherein coating the electrode with a dielectric material in circumferential segments along the length of the electrode.

14. The method of claim 9, wherein the electrode is a helical coil, and decreasing the surface area of the electrode includes increasing the pitch of the helical coil of the electrode.

15. The method of claim 9, wherein decreasing the surface area of the electrode includes reducing the cross-sectional dimensions of the electrode.

16. The method of claim 9, wherein decreasing the surface area of the electrode includes reducing the length of the electrode.

17. A system for treating the heart, coinprising:
a cardiac harness configured to conform generally to at least a portion of a patient's heart;
an electrode associated with the cardiac harness and positioned on or proximate to the epicardial surface of the heart, the electrode having a pericardial side opposite an epicardial side;
a power source in communication with the electrode, the electrode and power source are at least a part of an electrical circuit; and an insulation disposed on the pericardial side of the electrode, wherein the impedance of the electrical circuit is greater than about 10 ohms.

18. The system of claim 17, wherein the impedance of the electrical circuit is greater than about 20 ohms.

19. The system of claim 17, further comprising a conductor in communication with the electrode and the power source, wherein the conductor includes less than about 50% silver.

20. The system of claim 17, wherein the insulation is a dielectric material.