HAVING A SMALLER DISPLACEMENT VOLUME

TECHNICAL FIELD

The present invention relates generally to the field of automatic, implantable cardioverters and defibrillators. More particularly, the present invention relates to an implantable cardioverter defibrillator (ICD) that is a capacitor-discharge device having its internal components, including a battery and a capacitor, selected and arranged in such a manner that the ICD has a relatively smaller displacement volume that permits effective subcutaneous implantation of the device in the pectoral region of human patients.

BACKGROUND OF THE INVENTION

Existing implantable cardioverter defibrillators (ICDs) are typified by a relatively large size that usually requires implantation of the prosthetic device in the abdominal cavity of a human patient. In order to allow for effective subcutaneous implantation of a prosthetic device in the pectoral region of a human patient, the maximum size of the prosthetic device needs to be less than about 40-90cc, depending upon the physical size and weight of the patient. Unfortunately, all existing ICDs have total displacement volumes of at least HOcc or greater. Even though there are numerous advantages to developing an ICD having a displacement volume small enough to permit implantation of the device in the pectoral

region of a human patient, to date it has been difficult to develop a practical ICD having a total displacement volume of less than about lOOcc. For reasons of simplicity and compactness, existing ICDs are universally capacitor-discharge systems that generate high energy car dioversion/ defibrillation countershocks by using a low voltage battery to charge a capacitor over a relatively long time period (i.e., seconds) with the required energy for the defibrillation countershock. Once charged, the capacitor is then discharged for a relatively short, truncated time period (i.e., milliseconds) at a relatively high discharge voltage to create the defibrillation countershock that is delivered through implantable electrode leads to the heart muscle of the human patient.

One of the primary reasons why capacitor-discharge ICDs of a smaller volume have not been developed to date relates to the electrical requirements for storing the high energy cardioversion/ defibrillation countershocks that are currently used to defibrillate human patients. Cardioversion countershocks have delivered energies ^{^} of between about 0.5 to 5.0 Joules and are used to correct detected arrhythmias, such as tachycardia, before the onset of fibrillation. Defibrillation countershocks, on the other hand, have delivered energies greater than about 3.0 Joules and are use to correct ventricular fibrillation or an advanced arrhythmia condition that has not responded to cardioversion therapy.

Presently, all capacitor-discharge ICDs are designed such that the capacitor can store a maximum electrical charge energy of at least about 35 Joules. In contrast, implantable pacemakers, which currently have displacement volumes of less than 50cc, are designed to deliver pacing pulses of no more than about 50 μjoules. The requirement that a capacitor- discharge ICD be capable of storing an electrical charge with enough energy to deliver an electrical pulse almost one million times as large as that of an implantable pacemaker significantly increases the size of the ICD over the size of the pacemaker due to the size of the electrical components necessary to store this amount of electrical charge energy.

The accepted requirement that ICDs be capable of storing a

maximum electrical charge energy of at least about 35 Joules arises out of the definition of an appropriate safety margin for the device according to a clinically developed defibrillation success curve as shown in Fig. 11. The defibrillation success curve plots the percentage probability of successful defibrillation for a ventricular fibrillation of about 5-10 seconds versus the energy of a monophasic defibrillation countershock as measured in Joules. The safety margin for a given device for a given patient is presently accepted to be the difference between the maximum electrical charge energy (E _c ) stored by the capacitor in that device and the median defibrillation threshold energy (DFT) required for that patient.

Under existing medical practice, each time an ICD is implanted in a human patient, an intraoperative testing procedure is attempted in order to determine the median DFT for that patient for the particular electrode lead combination which has been implanted in the patient. The intraoperative testing procedure involves inducing ventricular fibrillation in the heart and then immediately delivering a defibrillation countershock through the implanted electrode leads o a specified initial threshold energy, for example, 20 Joules for a monophasic countershock. If defibrillation is successful, a recovery period is provided for the patient and the procedure is usually repeated a small number of times using successively lower threshold energies until the defibrillation countershock is not successful or the threshold energy is lower than about 10 Joules. If defibrillation is not successful, subsequent countershocks of 35 Joules or more are immediately delivered to resuscitate the patient. After a recovery period, the procedure is repeated using a higher initial threshold energy, for example, 25 Joules. It is also possible that during the recovery period prior to attempting a higher initial threshold energy, the electrophysiologist may attempt to lower the DFT for that patient by moving or changing the electrode leads. The intraoperative testing procedure is designed to accomplish a number of objectives, including patient screening and establishing a minimum DFT for that patient. Typically, if more than 30-35 Joules are

required for successful defibrillation with a monophasic countershock, the patient is not considered to be a good candidate for an ICD. Otherwise, the lowest energy countershock that results in successful defibrillation is considered to be the median DFT for that patient. The use of the lowest energy possible for a defibrillation countershock is premised on the accepted guideline that a countershock which can defibrillate at a lower energy decreases the likelihood of damage to the myocardial tissue of the heart. For a background on current intraoperative testing procedures, reference is made to M. Block, et al., "Intraoperative Testing for Defibrillator Implantation", Chpt. 3; and J.M. Almendral, et al., "Intraoperative Testing for Defibrillator Implantation", Chpt. 4, Practical Aspects of Staged Therapy Defibrillators. edited by Kappenberger, L.J. and Lindemans, F.W., Futura Publ. Inc., Mount Kisco, N.Y. (1992), pgs. 11-21.

Once the median DFT for a patient is established, the electrophysiologist will determine a safety margin for a given ICD device usually by subtracting the median DFT from the maximum E _c stored by that device. Alternatively, a different calculation for the safety margin is sometimes determined by estimating that point on the defibrillation success curve where the electrical energy of a defibrillation countershock will insure a 99% success (E ₉₉ ). Under either definition, the safety margin needs to be large enough to accommodate upward deviations along the defibrillation success curve. Such deviations may be expected, for example, with subsequent rescue defibrillation countershocks delivered later in a treatment after initial cardioversion or defibrillation countershocks of lesser energies were not successful. In these situations clinical data has found that, when delivered after 30 to 40 seconds of ventricular fibrillation, the electrical energy necessary to achieve effective defibrillation may increase 50% or more over the median DFT. As a result, an electrophysiologist usually will require that a given ICD have a first type of safety margin that is typically a factor of at least 2 to 2.5 times the median DFT for that patient before the electrophysiologist will consider implanting the given ICD in that patient. For the alternate E ₉₉

point safety margin, the electrophysiologist will require that a given ICD have a maximum E _c at least 10 Joules above the E ₉₉ point.

Based on current clinical data that the average median DFT is somewhere between 10-20 Joules for a monophasic countershock, the lower limit for the maximum E _c that must be stored by the ICD is accepted to be at least about 35 Joules, and more typically about 39 Joules, in order to generate a maximum defibrillation countershock having an adequate safety margin. The accepted lower limit for the maximum E _c of at least 35 Joules is supported by clinical evaluations, such as Echt, D.S., et al., "Clinical Experience, Complications, and Survival in 70 Patients with the Automatic Implantable Cardioverter/Defibrillator", Circulation. Vol. 71, No. 2:289-296, Feb. 1985. In this article, the authors evaluated data for early AICD devices having maximum E _c energies of 32 Joules stored in a 120 μF capacitor with a discharge voltage V _d of 750 Volts. In analyzing the clinical data for minimum DFTs, the authors concluded that the 32 Joule device had insufficient energy for effective defibrillation. It should be noted that in the next generation of the particular AICD devices studied, the maximum E _c for the device (the CPI Ventak®) was increased to 39.4 Joules by increasing the capacitance value of the ICD by using a 140 μF capacitor. Unfortunately, the requirement that an ICD be capable of storing a maximum E _c of this magnitude effectively dictates that the size of the ICD be greater than about lOOcc. This relationship between the maximum E _c that is required for an ICD and the overall size of the ICD can be understood by examining how an ICD stores the electrical energy necessary to deliver a maximum defibrillation countershock.

The only two components that impact on the ability of a capacitor- discharge ICD to store a maximum E _c are the capacitor and the battery, which together occupy more than 60% of the total displacement volume of existing ICDs. Thus, it will be apparent that the size of a capacitor- discharge ICD is primarily a function of the size of the capacitor and the size of the battery. For a capacitor, the physical size of that capacitor is principally determined by its capacitance and voltage ratings. The higher

the capacitance value, the larger the capacitor. Similarly, the physical size of a battery is also principally determined by its total energy storage, as expressed in terms of Amp-hours, for example. Again, the higher the Amp-hours, the larger the battery. With these concepts in mind, it is possible to evaluate how a maximum E _c affects the size of the capacitor and the size of the battery in an ICD.

The maximum electrical charge energy (E _c ) of an ICD is usually defined in terms of the capacitance value (C) of the capacitor that stores the charge and the discharge voltage (V _d ) at which the electrical charge is delivered as defined by the equation:

E _c = 0.5 * C * V _d 2 (Eq. 1)

The maximum electrical charge energy (E _c ) can also be defined in terms of how the energy is transferred from the battery to the capacitor. In this case, E _c is determined by the charging efficiency (e _c ) of the circuitry charging the capacitor, the battery voltage (V _b ), the battery current (l _b ) and the charging time (tc) as defined by the equation:

Ec = e _c * V _b * lb * tc (Eq. 2)

When Eqs. 1 and 2 are used to calculate a maximum E _c to be stored by the device, the capacitance value (C) and the charging time (t _c ) end up being the only true variables in these equations because the remaining values are all effectively determined by other constraints. In Eq. 1, for example, the discharge voltage (V _d ) for present ICDs can be no more than about 800 Volts due to voltage breakdown limitations of high power microelectronic switching components. As a result, V is typically between 650-750 Volts. In Eq. 2, it will be found that, for batteries suitable for use in an ICD, the maximum battery output voltage (V _b ) for ICDs is typically less than 6 Volts and, due to internal impedances within these batteries, the maximum battery current (l _b ) is about 1 Amp. In addition, the charging efficiencies (e _c ) of existing ICDs are presently on the order of about 50%.

When Eqs. 1 and 2 are evaluated for any given maximum E _c , it will be found that there necessarily is a minimum capacitance value (C _m i _n ) for

the capacitor and a minimum charging time (t _m i _n ) required to store that maximum E _c in the capacitor of the ICD. Knowing E _c and Vd, Eq. 1 can be reworked as follows to solve for C _m i _n :

O _n * = 2E _C /V _d 2 (Eq. 3) = 2 * 35 Joules / ( 750 Volts )2

= 124.4μF Similarly, knowing E _c , V _b , l _b and e, Eq. 2 can be reworked as follows to solve for t _m i _n : tmin = E _c / (e * V _b * lb) (Eq.4) = (35 Joules) / ( (0.50) * (6 Volts) * (1 Amp) )

= 12 seconds

In other words, the fact that all ICDs presently use a maximum E _c of at least 35 Joules means that all existing ICDs will require capacitors of greater than 124 μF, and that all existing ICDs which draw 1 Amp of current from the battery will have a charging time of greater than 12 seconds. Because the physical size of the capacitor is directly proportional to the capacitance rating of the capacitor in farads for a fixed voltage, the requirement that the capacitor be at least 124 μF is effectively a minimum size limitation on the capacitor for discharge voltages of less than about 800 Volts. Similarly, the requirement that each charging time for a defibrillation countershock draw at least 12 Amp-se nds of current from the battery is also a constructive minimum size limitation on the battery.

Thus, it can be seen that the existing requirement for a maximum E _c of at least about 35 Joules effectively dictates the size of both the capacitor and the battery and, consequently, the size of the ICD.

While existing ICDs have been successful in defibrillating human patients, and thereby saving lives, these devices are primarily limited to implantation in the abdominal cavity due to their relatively large size of greater than llOcc. It has long been recognized that it would be advantageous to reduce the total displacement volume of an ICD sufficiently to allow for subcutaneous implantation of the device in the pectoral region of human patients. This can only be done, however, so

long as the device provides for a sufficient safety margin to insure its effectiveness. Accordingly, it would be desirable to provide for an arrangement and configuration of the internal components of a capacitor- discharge ICD such that the total displacement volume of the ICD is reduced, while a sufficient safety margin for the device is retained.

SUMMARY OF THE INVENTION

The present invention is a capacitor-discharge implantable cardioverter defibrillator (ICD) having a relatively smaller displacement volume of less than about 90cc that permits effective subcutaneous implantation of the device in the pectoral region of human patients. The smaller volume of the ICD of the present invention is achieved by selecting and arranging the internal components of the capacitor-discharge ICD in such a manner that the ICD delivers a maximum defibrillation countershock optimized in terms of a minimum physiologically effective current (I _pe ) _/ rather than a minimum defibrillation threshold energy (DFT). One of the important results of optimizing the maximum defibrillation countershock in terms of a minimum effective current I _pe is that there is a significant decrease in the maximum electrical charge energy (E _c ) that must be stored by the capacitor of the ICD to less than about 30 Joules, even though a higher safety margin is provided for by the ICD. Due to this decrease in the maximum E _c , as well as corollary decreases in the effective capacitance value required for the capacitor and the net energy storage required of the battery, the overall displacement volume of the ICD of the present invention is reduced to the point where subcutaneous implantation of the device in the pectoral region of human patients is practical.

By using a physiologically effective current (I _pe ) to determine what is a safe and effective maximum defibrillation countershock, the present invention takes advantages of the realization that it is the effective current delivered to the heart by the defibrillation countershock, and not the total energy of the defibrillation countershock, that results in effective

defibrillation. In other /ords, the present invention recognizes that all Joules are not created equal and that the cells in the heart muscle will make more effective use of some types of electrical energy and less effective use of other types of electrical energy. The prior art technique of using a minimum DFT energy of the defibrillation countershock to establish safety margins effectively ignores the accepted fact that defibrillation countershock waveforms which differ in shape, tilt and duration, for example, can have significantly different defibrillation threshold energies. In contrast, the effective current I _pe as used by the present invention automatically compensates for any differences in the effectiveness of different waveforms. Consequently, the ICD of the present invention uses a minimum effective current I _pe delivered to the heart muscle, rather than using a minimum DFT energy, as the measure for insuring an adequate safety margin for the device. In accordance with a first aspect of the present invention, an implantable cardioverter defibrillator for subcutaneous positioning within a human patient comprises: a sealed housing structure constructed of a biocompatible material and having a displacement volume of less than about 90cc; one or more connector port means, each connector port means disposed in a wall of said housing structure for providing electrical connections between an interior space of said housing structure and a corresponding electrode lead within said human patient; circuit means disposed within said interior space of said housing structure and operably connected to said connector port means for sensing cardiac signals received from one or more of said electrode leads and, in response to the detection of an arrhythmia in said cardiac signals, controlling delivery of one or more high energy electrical cardioversion/ defibrillation countershocks of at least 0.5 Joules to the myocardium of said human patient; capacitor means disposed within said interior space of said housing structure and operably connected to said circuit means for storing electrical energy to generate said electrical cardioversion/ defibrillation countershocks and having an effective capacitance value of less than

120μF; and battery means disposed within said interior space of said housing structure and operably connected to said circuit means and said capacitor means for providing electrical energy to said circuit means and said capacitor means, wherein said battery means and said capacitor means are selected such that a maximum electrical charge energy stored by said capacitor means for each of said electrical cardioversion /defibrillation countershocks is less than about 30J and said implantable cardioverter defibrillator is capable of delivering at least five of said electrical cardioversion/ defibrillation countershocks in a four minute period. In accordance with a second of the present invention, the same elements as the first aspect are recited wherein said battery means and said capacitor means are selected to maximize a physiologically effective current (I _pe ) for a maximum electrical charge energy (E _c ) stored by said capacitor means. In accordance with a third aspect of the present invention, an implantable cardioverter defibrillator for subcutaneous positioning within the pectoral region of a human patient comprises: a sealed housing structure constructed of a biocompatible material and having a displacement volume of less than about 90cc and including one or more connector ports disposed in a wall of said structure for providing electrical connections between an interior space of said structure and electrode leads in said patient; circuit means within said interior space for sensing cardiac signals received from said electrode leads and, in response to the detection of an arrhythmia in said cardiac signals, controlling delivery of one or more electrical cardioversion/ defibrillation countershocks of at least 0.5 Joules to said patient; capacitor means within said interior space for storing electrical energy to generate said electrical cardioversion/ defibrillation countershocks and having an effective capacitance of less than about 120μF; and battery means within said interior space for providing electrical energy to said circuit means and said capacitor means and capable of charging said capacitor means in less than about 10 seconds to a maximum electrical charge energy of less than about 30 Joules.

In accordance with a fourth aspect of the present invention, an implantable cardioverter defibrillator for subcutaneous positioning within the pectoral region of a human patient diaving electrical energy storage requirements selected so as to treat mild cardiac disrhythmia conditions comprises: a sealed housing structure constructed of a biocompatible material and having a displacement volume of less than about 50cc and including one or more connector ports disposed in a wall of said structure for providing electrical connections between an interior space of said structure and electrode leads in said patient; circuit means within said interior space for sensing cardiac signals received from said electrode leads and, in response to the detection of an arrhythmia in said cardiac signals, controlling delivery of one or more electrical cardioversion/ defibrillation countershocks of at least 0.5 Joules to said patient; capacitor means within said interior space for storing electrical energy to generate said electrical cardioversion/ defibrillation countershocks and having an effective capacitance of less than about 80μF; and battery means within said interior space for providing electrical energy to said circuit means and said capacitor means and capable of charging said capacitor means in less than about 10 seconds to a maximum electrical charge energy of less than about 25 Joules, wherein the total amount of electrical energy stored by said battery means is less than 12,000 Joules and the budgeted number of electrical cardioversion/ defibrillation countershocks is less than about 200.

To understand how the present invention can use a minimum effective current I _pe to insure an appropriate safety margin for the ICD, it is necessary to recognize that the objective of any defibrillation countershock is to generate an electric field across as much as possible of the heart muscle, the myocardium. This electric field must have a current strong enough to extinguish all cardiac depolarization wavefronts in the myocardium, and the current must be strong enough to prevent the myocardium cells from being restimulated during their vulnerable period. In essence, the present invention recognizes that the electric current generated by the defibrillation countershock must be larger than whatever

minimum electric current is required for cell stimulation by at least a sufficiency ratio that will insure successful defibrillation. In this way, the use of an effective current I _pe can be thought of as a correction factor applied to the actual current of the defibrillation countershock in order to compensate for the cellular phenomenon that currents below some minimum value simply do not have any effect on the cells.

It has long been known that in order to stimulate cells, a current applied to those cells must have a value at least equal to a rheobase value of those cells, otherwise the current applied to the cells is not effective in stimulating the cells. G. Wiess, "Sur la Possibilite de Rendre Comparable entre Eux les Appareils Suivant a l\'Excitation Electrique", Arch. Ital. deBiol. Vol. 35, p. 41 (1901); and L. Lapicque, "Definition Experimetelle de l\'excitabilite", Proc. Soc. deBiol.. Vol. 77, p. 280 (1909). Lapicque defined the rheobase value as the stimulating current required for a pulse of infinite duration. From this definition, he further defined a chronaxie value (d _c ) to be the duration of a pulse that required a current twice that of the rheobase value. These two works have been combined in the literature to define a strength-duration model for the required average current for neural stimulation known as the Weiss-Lapicque strength-duration curve, an example of which is shown in Fig. 12.

The present invention builds on the Weiss-Lapicque strength- duration model to define a physiologically effective current I _pe as a simple model for the efficiency of a monophasic defibrillation countershock in terms of the actual average current of the defibrillation countershock. The actual average current (I _a v _e ) is given by the amount of electrical charge delivered at the electrode leads divided by the duration of the pulse delivering that charge. The end result of the derivation of a definition of effective current I _pe as taught by the present invention is that the effective current I _pe is given by the charge delivered to the electrode leads divided by the sum of the pulse duration (d) and the chronaxie time constant for the heart (d _c ). Expressing the charge delivered to the electrode leads in terms of the actual average current I _aV e yields a definition equation as

follows:

I _pe = (I _ave * d) / (d + d _c ) (Eq.5)

It can be seen from Eq. 5 that if the chronaxie value d _c were zero, the effective current I _pe would simply be I _aV e, the average current of a monophasic defibrillation countershock. In this way, the definition of an effective current I _pe distills the information contained in the Weiss- Lapicque strength duration curve to correct the actual average current I _aV e of a monophasic defibrillation countershock in order to compensate for the chronaxie phenomenon of the cells of the myocardium. When a minimum effective current I _pe is used to select and arrange the internal components of a capacitor-discharge ICD, the end result is a pair of surprising and non-intuitive conclusions.

First, the optimum capacitance value for the capacitor in a capacitor- discharge ICD is not determined by any stored or delivered energy requirement, but instead is a relatively constant value much smaller than any currently used capacitance values. The use of a minimum effective current I _pe predicts that the optimum capacitance value will be a function of only the chronaxie time constant and the inter-electrode resistance of the electrode leads. This means that a capacitor with a smaller effective capacitance actually delivers a defibrillation countershock with more effective current I _pe than a capacitor having a larger effective capacitance. When the optimum capacitance value is analyzed in terms of effective current I _pe , it is found that the optimal capacitance value is given by the formula: - C = ( 0.8 * dc ) / R (Eq. 6)

Second, there is no single optimum pulse duration for a defibrillation countershock having an arbitrary capacitance value. Instead, a defibrillation countershock of a shorter duration can provide a more effective current I _pe than a defibrillation countershock of a longer duration. The use of a minimum effective current I _pe predicts that the optimum pulse duration is a compromise between the RC time constant of the capacitor-discharge circuitry and the heart\'s defibrillation chronaxie

time constant, d _c . Thus, the predicted optimum pulse duration is not a constant, but rather is a function of the effective capacitance and other variables. The predicted optimum pulse duration can be most simply, and robustly, expressed as a fixed tilt or exponential decay followed by a fixed time duration extension. When the optimum pulse duration value is analyzed in terms of effective current I _pe , it is found that the optimal pulse duration is given by the formula: d = ( ( R * C ) + d _c ) / ( e - l) (Eq. 7)

Because the physical size of the capacitor is a function of its capacitance rating, the use of a capacitor with a smaller effective capacitance provides for a significant reduction in the displacement volume of the capacitor. In addition, because less energy is required to charge up a capacitor with a smaller effective capacitance, a battery with a smaller total energy storage, and, hence, a smaller displacement volume, may also be used. Finally, the shortening of the duration of the defibrillation countershock further decreases the energy requirements of both the capacitor and the battery, and also improves the safety margin of the device. In the preferred embodiment, several additional innovations are also used to further enhance the effectiveness of the defibrillation countershock and decrease the energy storage requirements of the ICD.

As a result of all of these improvements in the selection and arrangement of the internal components of the ICD of the present invention, the capacitor in the device only needs to store a maximum E _c of less than about 30 Joules, and preferably less than 27 Joules. The effective capacitance of the capacitor required by the present invention can be less than 120 μF, and preferably less than about 95 μF. By optimizing both the charging time and the countershock duration for the smaller maximum E _c , the size of the battery required by the present invention is reduced because the total energy storage capacity of the device can be less than about 1.0 Amp-hours. In the preferred embodiment, the charging time for each defibrillation countershock is reduced to less than about 10 seconds and the pulse duration of a monophasic defibrillation

countershock, or of a first phase of a multiphasic defibrillation countershock, is reduced to less than about 6 milliseconds.

By significantly reducing the displacement volume of both the capacitor and the battery, the overall displacement volume of an ICD in accordance with the present invention can be reduced below 90cc, and preferably to between 40-80cc, and with current capacitor and battery technology between about 50-70cc. Because the size requirements for effective pectoral implantation will be distributed across the range from 40-

90cc for the entire population, it is obvious that the smaller the overall displacement of the ICD, the greater the percentage of human patients who can benefit from pectoral implantation of the device. At the displacement volumes provided for by the present invention, subcutaneous implantation of the device in the pectoral region of a human patient can be quite practical and effective.

BRIEF DESCRIPTION OF THE DRAWINGS Fig. 1 is a frontal plan view showing the automatic, implantable cardioverter defibrillator of this invention implanted in the pectoral position of a human patient. Figs. 2 and 3 are frontal and side plan views, respectively, of the preferred embodiment of the ICD of the present invention.

Figs. 4 and 5 are side and frontal plan views, respectively, showing the power, capacitor, circuit and connector ports means positioned in the preferred embodiment of the ICD of the present invention. Figs. 6 and 7 are plan views, showing the interior of the preferred embodiment of the ICD of the present invention.

Fig. 8 is a voltage versus time graph showing the relative distribution of the defibrillation energy discharged from the capacitor of the preferred embodiment of the ICD of the present invention. Figs. 9 and 10 are voltage versus time graphs and a comparison table, respectively, showing the defibrillation energy discharged from the capacitor used in the preferred embodiment of the ICD of the present

invention versus the defibrillation energy discharged from a capacitor in a prior art ICD.

Fig. 11 is a defibrillation success curve used to define a minimum defibrillation threshold (DFT) for prior art ICDs for monophasic intravenous defibrillation countershocks.

Fig. 12 is a typical Weiss-Lapicque strength-duration curve showing the average current required for defibrillation as a function of the pulse duration.

Fig. 13 is a defibrillation success curve for the present invention using an physiologically effective current (I _pe ) for monophasic intravenous defibrillation countershocks.

Fig. 14 is a graph of the minimum effective current (I _pe ) for monophasic intravenous defibrillation countershocks versus fibrillation time showing the impact of prolonged fibrillation on minimum I _pe . Fig - 15 is a graph of the minimum effective current (I _pe ) for monophasic intravenous defibrillation countershocks as a function of electrode resistance for both fixed duration and tilt countershock pulses.

Fig. 16 is a block diagram of a dual battery system energy storage system for the preferred embodiment of the present invention. Fig. 17 is a block diagram of a rechargeable version of the dual battery system shown in Fig. 16.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

The following detailed description, when considered in connection with the accompanying drawings in which like reference numerals des¬ ignate like parts throughout the figures, describes some of the embodiments of the present invention. In describing the present invention, first a description of the preferred mechanical arrangement of the internal components of the implantable cardioverter defibrillator (ICD) will be presented to provide a context for the remainder of the description. Next, a mathematical explanation of the derivation of the physiologically effective current (I _pe ) as used by the present invention will

be presented. Then, each of the major features responsible for decreasing the overall displacement volume of the ICD will be described. These features include: the use of a more optimal pulse duration and pulse waveform for the cardioversion/ defibrillation countershock, the use of a capacitor having a smaller effective capacitance value, and the use of an improved battery configuration having a smaller total energy storage.

Mechanical Arrangement of the ICD Components

Fig. 1 shows an automatic ICD 17 of the present invention implanted in the pectoral region 18 of the chest 11 of patient 10. The ICD 17 has a plurality of connector ports for connection to various implantable catheter and other electrode means, as is known in the art. For example, electrode leads 41 and 42 are shown extending form the ICD 17 to catheter electrodes 40 and 15 which are passed, respectively, into the superior vena cava 14 and the right ventricle 13 of heart 12. Further, lead 43 is shown extending from the ICD 17 to a subcutaneous patch electrode 16. The specific configuration of the electrodes of the defibrillation system is dependent upon the requirements of the patient as determined by the physician. Figs. 2 and 3 show the ICD 17 comprised of a housing 19 having mating half shells 21 and 22. Positioned and mounted on top of housing

19 is a top connector portion 20 having a plurality of connecting ports 23 which are described further below. Importantly, the ICD 17 is comprised of a compact, self contained structure having predetermined dimensions which permits pectoral implantation. The housing 19 and top connector

20 are constructed and arranged to yield a cooperating structure which houses power means, control means and capacitive means. This cooperating structure permits subcutaneous implantation in the pectoral region of a human patient and provides a compact and effective ICD that automatically senses the bioelectrical signals of the heart and is able to provide a 750 volt capacitive discharge, for example, to the heart for defibrillation purposes.

In the past, ICDs have required a size and configuration for functional purposes that necessitated implantation in the abdominal cavity of a patient. Such implantation has resulted in patient discomfort. However, the physical parameters of these prior art devices have prevented pectoral implantation, which is preferred by physicians and patients alike. Table 1 below shows the size and weight comparisons between known prior art ICD devices and the ICD 17 of the present invention.

TABLE 1

(120CC) (60CC)

As shown in Table 1, the ICD 17 of this invention, provides a structure which is 50% of the volume of prior art devices and which has a weight which is 55% of the weight of the prior art devices. The connector port, capacitor, battery and electronic circuitry of the ICD 17 of the present invention are further described below.

It is important in this invention that the ICD 17 be constructed and arranged to minimize the overall displacement volume of the device to allow for pectoral implantation, for example. The housing structure -19 is a compact and lightweight structure made of a biocompatable material and

has a contoured configuration. The overall structure of this invention has a weight of less than 130 grams, and preferably less than 120 grams, and a volume of less than 90cc, and preferably between about 40-80cc. As shown in Table 1, the ICD 17 of this invention has generally 55% of the weight of prior art devices and a volume which is generally 50% of that of prior art devices. Table 1 further shows the weights and volumes of the respective components of this invention (connector, capacitor, batteries and electronics) as a percentage in weight and volume of the total and in comparison to prior art devices. As further shown in Figs. 2 and 3, the housing structure 19 has a contoured periphery which is matingly connected to the top connector member 20 which also has a mating contoured configuration. The housing 19 is constructed of a biocompatable material such as a titanium or a stainless steel alloy. The top connector member 20 is also constructed of a biocompatable material, such as a biocompatable polymeric composition. It has further been found that for pectoral implantation purposes, that the housing structure 19 have a desired length to width to thickness ratio of approximately 5 to 3 to 1.

When selected in accordance with the optimized minimum physiological current (I _pe ) as described below, the capacitor has an effective capacitance of approximately 85 uF, is constructed and arranged to deliver an initial discharge voltage V _d of 750 Volts, yielding the effective defibrillation countershock which is also described below. In the preferred embodiment, the effective discharge voltage and capacitance is achieved by using two flash-type capacitors in series, each having a capacitance rating of 170 μF and a voltage rating of 375 Volts, while occupying a total displacement volume of only 7cc each. The output of the capacitors is in communication with an electronic circuitry output portion that generally is comprised of a flash type circuit which delivers the capacitor discharge through electrodes 15, 16 and 40, for example.

Figs. 4 and 5 show the canister housing 19 having an interior space 30 wherein capacitors 26 and 27 are positioned and wherein a battery

system 28 and circuit board portions 31 and 32 are positioned. The top connector 20 is shown mounted to the top of the canister housing 19. Connecting ports 36, 37 and 39 are shown positioned in the top connector 20. The connector ports 36 and 37 are connectible to the positive defibrillating electrode, for example, while connecting port 38 is connectible to the negative defibrillating electrode, for example, and the connecting port 39 receives the pacing /sensing electrode leads 41, 42. Channels 24 and 25 provide communicative and fastener members that provide for the attachment of the top connector 20 to the canister housing 19 and for the electrical connection between the ports 36, 37, 38 and 39 and the electronic elements positioned in the interior space 30 of housing 19.

As discussed, the top connector 20 of the defibrillator ICD 17 has, for example, connecting ports 36 (DF+), 37(DF+), 38(DF-) and 39 (sensing/pacing). The lead connected to the DF- port, for example, is in conductive contact with the catheter electrode 15 placed in the right ventricle 13 of the heart 12. The electrode lead(s) connected to the DF+ port(s) are connected to either or both of the electrodes positioned in the superior vena cava 14 and the subcutaneous patch electrode 16. Alternatively, the DF+ port holes may not be utilized, and plugged by a stopper means, for example, when the ICD body itself is utilized as the positive element to complete the defibrillation circuit. The pacing/sensing electrode 44 provides an input to connecting port 39 of the ICD 17 and provides continual monitoring of cardiac signals from the heart. The circuitry of the ICD 17 has means to detect any tachycardiac or other arrhythmia condition and to thereby respond by the selective discharge of electrical energy stored in the capacitors 26 and 27.

As described in more detail below, the ICD 17 of this invention provides a device which utilizes smaller capacitors and batteries than those of prior art devices and thus yields a countershock generator device having a smaller displacement volume that permits effective implantation of the device in the pectoral region of a human patient. Although the smaller unit and associated components are smaller and

deliver a smaller energy countershock to the heart, the implantation of the device in the pectoral region provides for a better countershock vector. Together with the improved countershock pulse waveform as described below, the ICD 17 produces a more effective defibrillation/ cardioversion countershock than prior art ICD devices.

Figs. 6 and 7 show the mating housing half shells 21 and 22, respectively of canister housing 19. The half shell 22 is shown to have an interior peripheral band 34 which is fixed adjacent the peripheral edge 33. The interior peripheral band 34 extends outwardly from the edge 33 of half shell 22 and is constructed and arranged to receive the peripheral edge 35 of housing half shell 21. Alternatively, the peripheral band 34 may be mounted within housing half shell 21, whereby the half shell 22 is positioned thereabout. The peripheral band 34 is also provided to shield the electronic components within housing 19 during the welding process uniting the body shells 21 and 22.

The flexible circuit board 29 is mounted within the interior space 30 of housing 19. The circuit board 29 provides for the sensing/pacing circuitry in communication with the lead extending from connecting port 39, for example. When a fibrillation episode is detected, the circuit board 29 causes the capacitors 26, 27 to discharge an initial 750 Volt charge through the electrode leads connected to ports 36-38, for example, and to the heart 12 of the patient 10. The electronic circuitry has a sensing portion which monitors the heart beat rate irregularity by means of two small electrodes 44, as is known in the art. In the preferred embodiment, the circuitry further has a processor portion which determines, with respect to a predetermined standard, when the output portion of the circuit will be activated.

Fig. 8 is a graph showing the voltage discharge with respect to time from the 85 μF capacitor used in the preferred embodiment of the ICD 17 of this invention. The graph shows the incremental benefit of the voltage discharge with respect to time. Fig. 9 is a graph which shows the instantaneous voltage with respect to- time and compares the plotted

values of a countershock having the same delivered energy content for both the present invention and a typical prior art ICD. In Fig. 9, the countershock is a 20 Joule delivered energy monophasic countershock and it will be seen that the pulse duration of the countershock in accordance with the present invention is significantly shorter than the pulse duration of the countershock delivered by the prior art ICD.

As summarized in the table of Fig. 10, the 85 uF capacitor of the preferred embodiment of the present invention provides 25.33 Joules of delivered energy in the form of a biphasic defibrillation countershock having a delivery efficiency of 97.5% from a 26 Joule maximum E _c stored in the capacitors 26, 27. In comparison, the 140 uF capacitor used in a prior art ICD device provides 34 Joules of delivered in the form of a monophasic defibrillation having a delivery efficiency of 86.3% from a maximum E _c stored in the capacitor of 39.4 Joules. The effective current I _pe of the biphasic countershock delivered by the present invention is 5.67 Amps, uncorrected, and possibly as high as 6.55 to 7.32 Amps, when corrected to be a monophasic equivalent current. In contrast, the effective current I _pe of the monophasic countershock delivered by the prior art device is 6.79 Amps. Thus, the uncorrected I _pe of the present invention is only 20% less than the I _pe of the prior art device, while the maximum E _c of the present invention is more than 50% less than the maximum E _c of the prior art device.

When the corrected I _pe provided by the biphasic countershock of the preferred embodiment of the present invention is compared, the present invention provides essentially the same effective current I _pe as the prior art device with half the maximum E _c and, as little as half the requisite displacement volume for the capacitor. Depending upon the correction factor applied to convert the current efficiency of an optimized biphasic countershock pulse to a traditional monophasic countershock pulse (a 25% more energy efficient countershock is a 15% more efficient effective current, whereas a 40% more energy efficient countershock is a 28% more efficient effective current), the corrected I _pe of the present invention is

between 3% less to 7% more than the I _pe of the prior art device.

Derivation of the Physiologically Effective Current (l^Z)

The famous Weiss-Lapicque model was developed at the turn of the century. It was an empirical model and the first physiological explanation for why the model accurately predicts the required current for cellular stimulation was only recently explained. Irnich, W., "The Fundamental Law of Electrostimulation and its Application to Defibrillation", PACE 1990. 13 (Part 1): 1433-1447. The model gives the required (average) current for neural stimulation as:

I _aVe = K _{1 +} ( K ₂ / d ) (Eq. 8) with d being the pulse duration. The value Ki is the current required for an infinite duration pulse. The "chronaxie" is that duration which requires a doubling of the rheobase current. The chronaxie time constant d _c is thus given by: d _c = K ₂ / Kι (Eq. 9)

Defining I _r as the rheobase current gives: lave = I _r * [ l + ( c / d ) ] (Eq. 10)

The Weiss-Lapicque model was based on cell stimulation, not defibrillation. However, in 1978, Bourland et al showed, with a study of dogs and ponies, that defibrillation thresholds also followed the Weiss- Lapicque model when current averaged over pulse duration was used. Bourland, J.D., Tacker, W.A. and Geddes, L.A., "Strength Duration Curves for Trapezoidal Waveforms of Various Tilts for Transchest Defibrillation in Animals", Med. Instr.. (1978), Vol. 12, No. 1:38-41. A typical strength- duration curve is shown in Fig. 12.

Bourland et al. further proposed that the average current of a pulse was the best measure of its effectiveness when compared to other pulses of the same duration. This was found to hold fairly true for pulses from 2-20 ms in duration, regardless of waveform. Bourland, J.D., Tacker, W.A. and Geddes, L.A., et al. "Comparative Efficacy of Damped Sine Wave and Square Wave Current for Transchest Ventricular Defibrillation in

Animals", Med. Instr.. (1978), Vol. 12, No. 1:42-45.

Numerous studies have confirmed the strength-duration relationship for defibrillation currents. These same studies show that the defibrillation chronaxie time constant, d _c , is in the range of 2-4 ms. Using the available data on measured defibrillation chronaxie time constants, d _c

= 2.7 ± 0.9 ms is the average chronaxie value for the human heart.

In contrast to the accepted prior art technique of using a minimum defibrillation threshold energy (DFT) to measure the effectiveness of a defibrillation countershock, or even in contrast to the suggestion by Bourland et al to use the average current, the present invention defines an effective current as that percentage of the rheobase requirement for the human heart that the average current of a defibrillation countershock pulse can satisfy. Under this definition, successful defibrillation will require that I _aVe >= I _r * [ l + ( d _c / d ) ] (Eq. ll) where I _ave is the current averaged over the pulse duration of the defibrillation countershock. Satisfying this condition and substituting I _pe for I _r yields a definition of physiologically effective current (I _pe ) which can be expressed in several ways: Ip _e = I _ave / [ l + (d _c /d) ] (Eq. 12)

= (I _ave * d) / (d _c +d) (Eq. 13)

= delivered charge / (d _c +d) (Eq. 14)

Note that the effective current of a defibrillation countershock only equals the rheobase current when the output of the pulse is exactly operated at the defibrillation threshold, and, hence, with a zero safety margin. In general, the two parameters are not equal in value or orientation. The effective current I _pe is a system variable of the ICD, while the rheobase current I _r is primarily a physiologic variable.

Fig. 13 shows a defibrillation success curve for monophasic intravenous defibrillation countershocks plotted in terms of the effective current I _pe of the present invention. It will be apparent when comparing

the I _pe defibrillation success curve shown in Fig. 13 with the DFT defibrillation success curve shown in Fig. 11 that the I _pe curve is tighter and the necessary safety margin is much closer to the median I _pe required for effective defibrillation. Fig. 14 is a graph of the minimum I _pe for monophasic intravenous defibrillation countershocks versus fibrillation time showing the impact of prolonged fibrillation on minimum I _p e- Together, these figures illustrate how a device with a smaller maximum E _c can still provide a more than adequate safety margin, as long as the effective current I _pe of the defibrillation countershock is sufficient. The next two sections of the description set forth how to optimize the characteristics of a cardioversion/ defibrillation countershock in terms of effective current I _p e.

Optimal Pulse Waveform and Duration The present invention uses the effective current I _pe model to find the optimum pulse duration for a conventional, time-truncated, capacitor discharge defibrillation countershock waveform. In such a capacitor- discharge system, a capacitance (C) is ςharged to an initial voltage (Vj) and then discharged into the effective load resistance (R) of the heart for a pulse duration (d), at which time the capacitance will have a final voltage (V _f ). The amount of droop in the capacitor-discharge waveform at the time the pulse is truncated has been referred to as the "tilt" of the waveform as given by the equation:

Substituting the RC time constant exponential decay for V _f / Vj yields: tilt = 1 - e-d/RC (Eq. 16) where RC is the capacitor-discharge system time constant, also known as τ. Substituting the system parameters for C, V and tilt for the delivered charge in Eq. 14, the effective current I _pe can also be expressed as:

Ipe = ( C * V * tilt ) / ( d _c + d ) (Eq. 17)

= (C*V*(l-e-d/τ))/ (d _c + d) Because the first derivative of I _pe approaches zero at extreme values of d, its maximum is at the point of zero derivative. = CVi [ [ (d _c +d) [1/τ dΛ] - [1- e-dΛ] ] / (d _c +d)2 ]

= [(d+d _c )/τ]e iΛ-l+e- c )/τ] + l]e-d -l Normalizing Eq.18 to the system time constant by defining: z = d / τ (Eq.19) α = dc / τ (Eq.20)