CA2499842A1 - Apparatus and method for optimizing capacitor charge in a medical device - Google Patents
Apparatus and method for optimizing capacitor charge in a medical device Download PDFInfo
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- CA2499842A1 CA2499842A1 CA002499842A CA2499842A CA2499842A1 CA 2499842 A1 CA2499842 A1 CA 2499842A1 CA 002499842 A CA002499842 A CA 002499842A CA 2499842 A CA2499842 A CA 2499842A CA 2499842 A1 CA2499842 A1 CA 2499842A1
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
- A61N1/00—Electrotherapy; Circuits therefor
- A61N1/18—Applying electric currents by contact electrodes
- A61N1/32—Applying electric currents by contact electrodes alternating or intermittent currents
- A61N1/38—Applying electric currents by contact electrodes alternating or intermittent currents for producing shock effects
- A61N1/39—Heart defibrillators
- A61N1/3975—Power supply
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
- A61N1/00—Electrotherapy; Circuits therefor
- A61N1/18—Applying electric currents by contact electrodes
- A61N1/32—Applying electric currents by contact electrodes alternating or intermittent currents
- A61N1/38—Applying electric currents by contact electrodes alternating or intermittent currents for producing shock effects
- A61N1/39—Heart defibrillators
- A61N1/3975—Power supply
- A61N1/3981—High voltage charging circuitry
Abstract
A medical device for electrical termination of an arrhythmic condition of a patient's heart in embodiments of the invention may include one or more of t he following features: (a) at least one battery; (b) means for detection of an arrhythmic condition of a patient's heart; (c) at least one high voltage capacitor; (d) converter means for providing charging current from said battery to said capacitor; (e) means for maintenance of a charge on said capacitor between arrhythmia therapies; (f) controller means responsive to detection of an arrhythmic condition of said patient's heart and for providi ng a discharge control signal; and (g) discharge circuit means for delivering voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
Description
APPA RATUS AND METHOD FOR OPTIMIZING CAPACITOR CHARGE IN
A MEDICAL DEVICE
This application relates to the following concurrently filed, commonly assigned U.S. patent application: "Method and Apparatus for Maintaining Energy Storage in an Electrical Storage Device", reference number P-9171.00 filed September 30, 2002, which is incorporated herein by reference.
The present invention relates generally to stimulators for medical treatment by means of voltage shocks, and more particularly to cardioverters and defibrillators and electrode systems for use in conjunction therewith.
A defibrillator can be used to restore a normal heart rhythm by delivering an electrical shock to the heart when the heartbeat is dangerously fast due to ventricular tachycardia or ventricular ribrillation. Either of these conditions can reach a life-threatening point at which a pexson suddenly loses consciousness because the heart can no longer pump enough blood to meet the body's demand. For patients suffering from chronic arrhythmias involving ventricular tachycardia or ventricular fibrillation, a defibrillator can be surgically implanted in the patient's chest. The implanted defibrillator can be implanted into the chest of the patient during a minor surgical procedure.
An implantable cardioverter defibrillator (ICD) is a device that can be implanted in a patient's chest to monitor for and, if necessary, correct episodes of rapid heartbeat.
If the heartbeat gets too fast (ventricular tachycardia), the ICD can stimulate the heart to restore a normal rhythm. In cases where the heartbeat is so rapid that the heart cannot effectively pump any blood (ventriculax fibrillation), the ICD can provide an electric shock to "reset" the heartbeat.
The ICD gets its name from the two functions that it performs. First, the ICD
sends small electrical charges to the heart to "reset" it during ventricular tachycardia.
This process of converting one rhythm or electrical pattern to another is called cardioversion. Second, the ICD will send stronger charges to "reset" the heart if it begins ventricular fibrillation instead of beating. The act of stopping this potentially fatal quivering of the heart is called defibrillation. Although the main functions of the ICD are cardioversion and defibrillation, it can also be programmed to do anti-tachycardia and bradycardia pacing.
In anti-tachycardia pacing, when an ICD senses a fast but rhythmic heartbeat (tachycardia), it can release a series of low-intensity electrical pulses that gently interrupt the heart and allow it to return to a slower pace. In bradycardia pacing, when the ICD senses an abnormally slow heartbeat, it can send small electrical signals to pace the heart until it recovers and maintains a normal heart rate. These therapies are contrasted with both cardioversion and defibrillation, which involve high voltage shocks, which is the focus of the present invention.
In all of the ICD systems available today, a truncated capacitive-discharge shock is delivered by the ICD to electrodes that are positioned in, on, or near the heart. To generate the shock, existing ICD systems use an internal high current electrical battery cell connected to a step-up transformer and power conversion cixcuitry to charge one or more relatively small, but powerful, high voltage capacitors to provide a relatively high discharge voltage. When an electrical stimulation pulse is to be applied to the heart, the appropriate output switch is closed to connect the output capacitor to the cardiac tissue through the electrodes, thereby effectively "dumping" the charge stored in the output capacitor into the cardiac tissue. After the output decays to a predetermined output voltage, or after a predetermined shock duration has elapsed, the shock is truncated and the remaining energy in the output capacitor system is dissipated within the ICD system never being utilized or recovered.
The primary function of an ICD is to sense the occurrence of an arrhythmia, and to automatically apply an appropriate shock therapy to the heart aimed at terminating the arrhythmia. For example, if the ICD senses that the patient's heart is fibrillating then the ICD automatically delivers a high current shock to the patient's heart to defibrillate the organ. ICDs typically operate by first detecting the arrhythmia, then rapidly charging one or more storage capacitors contained within the device, and then quickly discharging the capacitors) to deliver the life saving shock therapy. However, a problem associated with rapidly charging a capacitor is that it creates a severe load on the battery. Thus reducing the battery's life.
An additional problem associated with the high voltage capacitors of an ICD is the amount of time it takes to charge the capacitors, typically about 5 to 20 seconds, Many studies have proposed that defibrillation and cardioversion shocks are most effective when delivered as quickly as possible following detection of arrhythmia. The chance of terminating an arrhythmia in a patient decreases as the length of time it takes for therapy to be delivered to the patient increases. Therefore, the shorter the charge time for the capacitors the more effective the defibrillation therapy.
Typically, ICD
battery sizes are proportional to the charging time. Therefore, the quicker the desired charging time, the larger the battery. In spite of this, it is desirable to make the ICD as small as possible and therefore large batteries are not desired and thus a balance must be struck between having a fast charging time and the size of the ICD.
Another problem involves providing a capacitor that maintains a high capacitance while at the same time has a reduced leakage current. The term "leakage current"
refers to the measure of stray direct current flowing through a capacitor after DC
voltage is impressed on it and is expressed in milliamps. The dielectric of a capacitor has a very high resistance, which prevents the flow of DC current. However there are some areas in the dielectric, which allow a small amount of current to pass.
The value of leakage current will continue to decrease while voltage is applied to the capacitor, until a very low steady state leakage current value is reached. However, as stated above, the present ICDs allow the remaining capacitor charge to dissipate after the arrhythmia has been treated. The longer capacitors are stored with no applied voltage, the higher the initial leakage current. Therefore, the constant recharging and the length between the recharging of the capacitors actually increases the amount of leakage current. A high leakage current can result in the poor performance and reliability of a capacitor. In particular, high leakage current results in a greater amount of charge leaking out of the capacitor once it has been charged. This is undesirable.
A MEDICAL DEVICE
This application relates to the following concurrently filed, commonly assigned U.S. patent application: "Method and Apparatus for Maintaining Energy Storage in an Electrical Storage Device", reference number P-9171.00 filed September 30, 2002, which is incorporated herein by reference.
The present invention relates generally to stimulators for medical treatment by means of voltage shocks, and more particularly to cardioverters and defibrillators and electrode systems for use in conjunction therewith.
A defibrillator can be used to restore a normal heart rhythm by delivering an electrical shock to the heart when the heartbeat is dangerously fast due to ventricular tachycardia or ventricular ribrillation. Either of these conditions can reach a life-threatening point at which a pexson suddenly loses consciousness because the heart can no longer pump enough blood to meet the body's demand. For patients suffering from chronic arrhythmias involving ventricular tachycardia or ventricular fibrillation, a defibrillator can be surgically implanted in the patient's chest. The implanted defibrillator can be implanted into the chest of the patient during a minor surgical procedure.
An implantable cardioverter defibrillator (ICD) is a device that can be implanted in a patient's chest to monitor for and, if necessary, correct episodes of rapid heartbeat.
If the heartbeat gets too fast (ventricular tachycardia), the ICD can stimulate the heart to restore a normal rhythm. In cases where the heartbeat is so rapid that the heart cannot effectively pump any blood (ventriculax fibrillation), the ICD can provide an electric shock to "reset" the heartbeat.
The ICD gets its name from the two functions that it performs. First, the ICD
sends small electrical charges to the heart to "reset" it during ventricular tachycardia.
This process of converting one rhythm or electrical pattern to another is called cardioversion. Second, the ICD will send stronger charges to "reset" the heart if it begins ventricular fibrillation instead of beating. The act of stopping this potentially fatal quivering of the heart is called defibrillation. Although the main functions of the ICD are cardioversion and defibrillation, it can also be programmed to do anti-tachycardia and bradycardia pacing.
In anti-tachycardia pacing, when an ICD senses a fast but rhythmic heartbeat (tachycardia), it can release a series of low-intensity electrical pulses that gently interrupt the heart and allow it to return to a slower pace. In bradycardia pacing, when the ICD senses an abnormally slow heartbeat, it can send small electrical signals to pace the heart until it recovers and maintains a normal heart rate. These therapies are contrasted with both cardioversion and defibrillation, which involve high voltage shocks, which is the focus of the present invention.
In all of the ICD systems available today, a truncated capacitive-discharge shock is delivered by the ICD to electrodes that are positioned in, on, or near the heart. To generate the shock, existing ICD systems use an internal high current electrical battery cell connected to a step-up transformer and power conversion cixcuitry to charge one or more relatively small, but powerful, high voltage capacitors to provide a relatively high discharge voltage. When an electrical stimulation pulse is to be applied to the heart, the appropriate output switch is closed to connect the output capacitor to the cardiac tissue through the electrodes, thereby effectively "dumping" the charge stored in the output capacitor into the cardiac tissue. After the output decays to a predetermined output voltage, or after a predetermined shock duration has elapsed, the shock is truncated and the remaining energy in the output capacitor system is dissipated within the ICD system never being utilized or recovered.
The primary function of an ICD is to sense the occurrence of an arrhythmia, and to automatically apply an appropriate shock therapy to the heart aimed at terminating the arrhythmia. For example, if the ICD senses that the patient's heart is fibrillating then the ICD automatically delivers a high current shock to the patient's heart to defibrillate the organ. ICDs typically operate by first detecting the arrhythmia, then rapidly charging one or more storage capacitors contained within the device, and then quickly discharging the capacitors) to deliver the life saving shock therapy. However, a problem associated with rapidly charging a capacitor is that it creates a severe load on the battery. Thus reducing the battery's life.
An additional problem associated with the high voltage capacitors of an ICD is the amount of time it takes to charge the capacitors, typically about 5 to 20 seconds, Many studies have proposed that defibrillation and cardioversion shocks are most effective when delivered as quickly as possible following detection of arrhythmia. The chance of terminating an arrhythmia in a patient decreases as the length of time it takes for therapy to be delivered to the patient increases. Therefore, the shorter the charge time for the capacitors the more effective the defibrillation therapy.
Typically, ICD
battery sizes are proportional to the charging time. Therefore, the quicker the desired charging time, the larger the battery. In spite of this, it is desirable to make the ICD as small as possible and therefore large batteries are not desired and thus a balance must be struck between having a fast charging time and the size of the ICD.
Another problem involves providing a capacitor that maintains a high capacitance while at the same time has a reduced leakage current. The term "leakage current"
refers to the measure of stray direct current flowing through a capacitor after DC
voltage is impressed on it and is expressed in milliamps. The dielectric of a capacitor has a very high resistance, which prevents the flow of DC current. However there are some areas in the dielectric, which allow a small amount of current to pass.
The value of leakage current will continue to decrease while voltage is applied to the capacitor, until a very low steady state leakage current value is reached. However, as stated above, the present ICDs allow the remaining capacitor charge to dissipate after the arrhythmia has been treated. The longer capacitors are stored with no applied voltage, the higher the initial leakage current. Therefore, the constant recharging and the length between the recharging of the capacitors actually increases the amount of leakage current. A high leakage current can result in the poor performance and reliability of a capacitor. In particular, high leakage current results in a greater amount of charge leaking out of the capacitor once it has been charged. This is undesirable.
Another problem associated with the present ICDs, is that the remaining charge after the arrhythmia is treated is just dissipated within the ICD. While the charge dissipated is relatively minimal when compared to the shock charge, after hundreds of shocks the remaining charges can add up to a substantial shock. Typically, 16 remaining charges can add up to provide a defibrillation shock. Further, the dissipated remaining charges equate to energy taken from the battery and never put to use.
Therefore, it would be desirable to capture these remaining charges and thus extend the Iife of the battery.
For the foregoing reasons, there is a need for an ICD, which allows for a relatively long charging time and yet retains clinical efficacy to prolong battery life and provide for a smaller battery. There is also a need for an ICD providing a high voltage capacitor with very low leakage current so that the capacitor could be held at full charge thus reducing the adverse effects of rapid charging. There is also a need for an ICD that when an arrhythmia is detected the ICD can deliver therapy at the quickest possible moment without having to wait for a capacitor to charge thus increasing the efficacy of the delivered therapy.
A medical device for electrical termination of an arrhythmic condition of a patient's heart in embodiments of the invention may include one or more of the following features: (a) at Ieast one battery; (b) means for detection of an arrhythmic condition of a patient's heart; (c) at least one high voltage capacitor; (d) converter means for providing charging current from said at least one battery to said at least one capacitor; (e) means for maintenance of a charge on said at least one capacitor between arrhythmia therapies; (f) controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and (g) discharge circuit means for delivering voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
A method for electrical termination of an arrhythmic condition of a patient's heart in embodiments of the invention may include one or more of the following features:
(a) charging at Ieast one high voltage capacitor with current from at least one battery, (b) detecting an arrhythmic condition of a patient's heart, (c) maintaining the charge on said at least one capacitor between arrhythmia therapies, (d) providing a controller means responsive to detection of an arrhythmic condition of said patient's heart, (e) generating a discharge control signal upon detection of an arrhythmic condition of said patient's heart; and (f) delivering a voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
FIG. 1 is a drawing illustrating the general physical components of a pacemaker/cardioverter/defibrillator and lead system of the type in which the present invention may be advantageously practiced;
FIG. 2 is a functional block diagram illustrating the general interconnection of voltage conversion circuitry of the present invention with the primary functional components of an implantable pacemaker/cardioverter/defibrillator;
FIG. 3 is a schematic block diagram of the general components of a pacemaker/cardioverter/defibrillator employing a high voltage charging circuit;
FIG. 4 is a flow diagram of an embodiment for capacitor optimization of the present invention;
FIG. 5 is a table representing a capacitor optimization embodiment of the present invention;
FIG. 6 is a flow diagram of an embodiment for capacitor optimization of the present invention;
The following detailed description is to be read with reference to the figures, in which like elements in different figures have like reference numerals. The figures, which are not necessarily to scale, depict selected embodiments and are not intended to limit the scope of the invention. Skilled artisans will recognize that the examples provided herein have many useful alternatives that fall within the scope of the invention.
The pxesent invention is not limited to implantable cardioverter defibrillators and may be employed in many various types of electronic and mechanical devices for treating patient medical conditions such as external cardioverter defibrillators, pacemakers, and neurostimulators. It is to be further understood; moreover, the present invention is not limited to medium current rate batteries and may be utilized for low and high current rate batteries. For purposes of illustration only, however, the present invention is below described in the context of medium current rate batteries and implantable cardioverter defibrillators.
The present invention is described generally in a system providing biphasic cardioversion pulses or shocks in a cardioversion system. However, it is fully contemplated that the present invention could be utilized in any type of pulse or shock delivery methodology utilizing any type of pulse of shock waveform without departing from the spirit of the invention. In the description of the preferred embodiment that follows, an implantable pacemaker/cardioverter/defibrillator in which the present invention is preferably implemented is capable of providing monophasic, biphasic, or any other caxdioversion pulse or shock waveform. However, a variety of implantable leads and electrode systems may be employed, with more than one cardioversion electrode connected electrically in common to widen the cardioversion energy distribution across the heart. Such electrodes may include indwelling right ventricular, superior vena cava, and coronary sinus electrodes, active pulse generator case electrodes and/or epicardial and subcutaneous patch electrodes in various combinations of two or more. With a three electrode system, two of the electrodes are connected in common, and the energy distribution between the two common and the third electrode may lead to reduced energy sufficient to reliably cardiovert a heart in fibrillation or high rate malignant ventricular tachycardia.
FIG. 1 illustrates such a general implementation of an implantable pacemaker/cardioverter/defibrillator 10 and one possible selection of cardioversion electrodes on associated electrical leads 14, 16 and 18, and their relationship to a human heart 12. The leads 14, 16, and 18 are coupled to the pacemaker/cardioverter/defibrillator 10 by means of a mufti-port connector block 20, which contains separate connector ports fox each of the three leads illustrated. Each of the leads 14, 16, 18 comprise a large surface area cardioversion electrode, and lead 18 _7_ also comprises a pair of pace/sense electrodes (making it a tripolar lead) all as described below.
Unipolar lead 14 is coupled to a subcutaneous cardioversion electrode 30, which is intended to be mounted subcutaneously in the region of the left chest.
Unipolar lead 16 is a coronary sinus (CS) lead employing an elongated coil, cardiovexsion electrode that is located in the coronary sinus of the heart. When positioned in the CS, the CS
electrode extends around the heart from a point within the opening or ostium of the CS
to a point in the vicinity of the left atrial appendage, as shown in broken line format at 32.
Tripolar lead 18 is provided with an elongated electrode coil 28 which is located in the right ventricle of the heart and functions as a third cardioversion electrode. Lead 18 also includes a first pace/sense electrode 34 and a second, closely spaced, pace/sense electrode 38. Electrode 34 takes the form of a distal helical coil, which is screwed into the myocardial tissue of the right ventricle. The second pace/sense electrode 38 is closely spaced to the electrode 34 for bipolar pacing and near fteld electrogram or R-wave sensing in the apex of the right ventricle. A more detailed description of the leads illustrated can be found in U.S. Pat. No. 5,163,427, herein incorporated by reference in its entirety.
Through testing at implantation of cardioversion efficacy across one of the three electrodes with the other two electrodes in common or with each of the other electrodes alone, a selection may be made of the most efficacious electrode selection.
If only two electrodes are needed, then the third lead and electrode may be eliminated.
Typically, it is expected that all thxee of the electrodes will be employed, with two connected electrically in common internally within the pulse generator 10 as described below.
FIG. 2 is a block diagram illustrating the general interconnections of a voltage output circuit 40, a voltage charging circuit 64 and capacitor bank 56, 58 according to one embodiment of the present invention with a prior art implantable pacemakex/cardioverter/defibrillator. As illustrated, the device is controlled by means _g_ of a stored program in a microprocessor 42, which performs all necessary computational functions within the device. Microprocessor 42 is linked to control circuitry 44 by means of a bi-directional datalcontrol bus 46, and thereby controls operation of the output circuitry 40 and the high voltage charging circuitry 64. On reprogramming of the device or on the occurrence of signals indicative of delivery of cardiac pacing pulses or of the occurrence of cardiac contractions, pace/sense circuitry 78 will awaken microprocessor 42 to perform any necessary mathematical calculations, to perform tachycardia and fibrillation detection procedures and to update the time intervals controlled by the timers in pace/sense circuitry 78.
The control circuitry 44 provides three signals of primary importance to the output circuitry 40 of the present invention. These include the first and second control signals discussed above, labeled here as ENAB, line 48, and ENBA, line 50, which govern the timing and duration of the two phases of the biphasic cardioversion pulse or shock.
Also of importance is the DUMP signal on line 52, which initiates discharge of the output capacitors, and the VCAP signal on line 54, which is indicative of the voltage stored on the output capacitors C1, C2, and is applied to the control circuitry 44.
As described above, a wide variety of cardioversion electrode bearing leads may be attached to two or all three cardioversion output terminals, labeled HVX, HVA, and HVB in FIG. 2, coupled to the connector block 20 bores. In the example illustrated in FIGS. 1 and 2, it will be assumed that the electrodes 28, 30 and 32 are coupled to the high voltage output circuitry 40 by means of connectors in the connector block illustrated as conductors 22, 24 and 26, respectively. As shown in FIG. 3, conductors 22 and 24 labeled HVX and HVA are electrically connected in common so that an output shock may be delivered even if all three leads 18, 14 and 16 and electrodes 28, 30 and 32, respectively, are connected to the pulse generator as shown in FIG.
1 and described above.
The high voltage output circuit 40 includes a capacitor bank, including capacitors 56 and 58, which is discussed in more detail below, and diodes 70 and 72, used for delivering defibrillation pulses to the electrodes. In FIG. 2, the capacitor bank is illustrated in conjunction with the high voltage charging circuitry 64, controlled by the control/timing circuitry 44 by means of CHDR line 66. As illustrated, capacitors 56 (C1) and 58 (C2) axe charged by means of a high frequency, high voltage transformer 68. Proper charging polarities are maintained by means of the diodes 70 and 72.
VCAP line 54 provides a signal indicative of the voltage on the capacitor bank, and allows for control of the high voltage charging circuitry and for termination of the charging function when the stored voltage equals the programmed charging level.
The delivery of the biphasic cardioversion shock is controlled by the partial discharge of the voltage on the output capacitor bank in a first direction during a first phase logic signal on ENAB, line 48, and by further discharge of the remaining voltage in a second direction during closely timed second signal on ENBA, line 50.
When ENAB is present, the first phase of the cardioversion pulse is delivered between the electrodes) 30 and/or 32 and electrode 28. During a logic signal on ENBA, line 50, the second phase is delivered between in the opposite direction between the same electrodes.
Pace/sense circuitry 78 includes an R-wave amplifier according to the prior art, or more advantageously as disclosed in U.S. Pat. No. 5,117,824 by Keimel et al, which is incorporated herein by reference in its entirety. However, the present invention is believed workable in the context of any known R-wave amplification system.
Pace/sense circuitry 78 also includes a pulse generator for generating cardiac pacing pulses, which may also correspond to any known cardiac pacemaker output circuitry and includes timing circuitry for def ning ventricular pacing intervals, refractory intervals and blanking intervals, under control of microprocessor 42 via control/data bus 80. Control signals triggering generation of cardiac pacing pulses by pace/sense circuitry 78 and signals indicative of the occurrence of R-waves, from pace/sense circuitry 78 are communicated to control circuitry 44 by means of a bi-directional data bus 81. Pace/sense circuitry 78 is coupled to helical electrode 34 and ring electrode 38 of tripolar lead 18 through connector elements of the connector block 20 and associated adapters, if necessary, illustrated schematically as conductors 36 and 37.
The present invention constitutes an apparatus and method for maintaining a full or partial charge on a capacitor within an implantable medical device between therapies.
The particular circuitry or components involved in the implementation of shock timing optimization axe shown in specific detail. However, it is fully contemplated that alternate circuitry or components could be utilized, such as described in U.S. Pat.
No. 6,438,420 (Thompson) herein incorporated by reference, without departing from the spirit of the invention. A number of additional expressions for input and output signals or terminals than those described above axe used throughout, including:
CHGDR--Charge drive signal for driving the on/off switch in the primary winding of the flyback transformer at a duty cycle established by the relative on and off times.
VSS--VSS is the circuit ground, which may also appear labeled QVSS and may be connected to BATTN.
BATT--Battery positive power supply, which may also appear as B+ or as BP.
BATTN--Battery negative power supply.
PPLUS--Plus terminal for the pace/sense function.
PMINUS--Negative terminal for the pace/sense function.
ENBA--Enable signal commanding capacitor discharge from HVB to HVA (and HVX) and setting the duration of one phase of the biphasic pulse.
ENAB--Enable signal commanding capacitor discharge from HVA (and HVX) to HVB and setting the duration of the other phase of the biphasic pulse.
CSP--Charge store positive terminal.
C1P--Capacitor 1 positive terminal connection.
C1N--Capacitor 1 negative terminal connection.
C2P--Capacitor 2 positive terminal connection.
C2N--Capacitor 2 negative terminal connection.
-lI-CSN--Charge store negative terminal.
VDD--Internally generated programmable regulated power supply.
DUMP--DUMP signal initiates the internal self discharge of the capacitoxs C1, to a load impedance.
OPTIN--Input terminal to the drive circuit optionally connected to an opto-coupler.
VIN--Input terminal to the drive circuit optionally connected to an input signal source.
VOUT--Output terminal of the drive circuit for supplying VDD voltage.
CSEN--Enable signal input terminal of the drive circuit optionally coupled to receive an opto-coupler command signal.
CSOUT--Output terminal of the drive circuit optionally coupled to drive an opto-coupler.
Other acronyms may appear in the description of the following drawings, which will be explained as necessary to understand the manner in which the present invention may be practiced in its preferred embodiment.
Turning now to FIG. 3, the circuit components of the pacemakerlcardioverter/defibrillator of the present invention are depicted and they include the batteries 11 and 13, the PC board 102, the high voltage output capacitors C1, C2 (56, 58 in FIG. 2), the high power hybrid board 104, the low power hybrid boaxd 106, the crystal 15, the antenna 17, and the reed switch 19. The batteries 11 and 13 are coupled to the BATT and BATTN inputs of the PC board 102. Although two batteries are shown, it is fully contemplated that any type or combination of batteries could be utilized, such as a single cell battery, a dual cell battery, or a mixture of high current and low current cells, without departing from the spirit of the invention. The crystal 15 is coupled to the Xl and X1N inputs of the low power hybrid 106.
The antenna 17 is coupled between the ANT and ANTGND inputs of low power hybrid 106 and the reed switch 19 is coupled between the RDSW and RSGND inputs of low power hybrid 106. The PPLUS and PMINLTS terminals are coupled to respectively labeled pins of the low power hybrid 106, which contains the pace/sense circuitry 78 of FIG. 2.
The low power hybrid 106 includes the basic timing and control circuitry of the system, including the programming and telemetry functions, the electrogram sensing and pacing functions, the microprocessor and RAM/ROM memories, all implemented in both digital and analog circuits corresponding to blocks 42, 44 and 78 in FIG. 2.
The low power hybrid 106 develops the CHGDR signal as well as the DUMP, ENBA
and ENAB signals relevant to the operation of the high voltage output circuit of the present invention.
The PC board 102 corresponds to the high voltage-charging block 64 in FIG. 2, and also includes the step up transformer 110 and diodes 121, 123. The relatively large output capacitors C1, C2 are electrically connected to the PC boaxd 102 through the input terminals C1N and C1P and C2N and C2P, respectively. The PC board presents the charge storage positive and negative signals CSP and CSN, respectively, to the high power hybrid 104. PC board 102 also includes an on-off control switch, responsive to the CHGDR signal from the low power hybrid 106, for supplying stepped up, rectified current to the output capacitors C1, C2, across which the voltage signals CSP, CSN are developed.
The high power hybrid 104 corresponds to the high voltage output block 40 illustrated in FIG. 2 and includes switching circuitry for delivery of voltage stored in capacitors C1 and C2 as monophasic, biphasic, or any other output pulse waveform.
Delivery of the output pulses is controlled by the low power hybxid 106 via ENAB and ENBA lines 48 and 50, respectively. Similarly, the HVA line 24, which is coupled in common to the HVX line 22, and the HVB line 26 are coupled to the HVA and HVB
output pins of high power hybrid 104. The high voltage discharges forming the cardioversion shocks are generated from the high power hybrid 104 and conducted to the HVA and HVB output terminals and the cardioversion electrode system employed as described above.
With reference to Figures 2 and 3 again, one embodiment of the present invention is described. In one embodiment, capacitors 56 and 58 are high voltage capacitors with an extremely low leakage current. An exemplary low leakage capacitor is described in U.S. Pat. No. 5,808,856 (Bischoff, et. al.), U.S. Pat. No.
6,426,863 (Munshi) and U.S. Pub. No. 2002/0052078 (Zheng et. al.). While it is preferable that a low leakage capacitor be utilized for the present embodiment, it is contemplated that any high voltage capacitor could be utilized without departing from the spirit of the invention. Further, it is fully contemplated that the present invention could utilize one or more individual capacitors as well as multiple capacitors utilizing a wide range of capacitor voltages. Nevertheless, preferably capacitors 56 and 58 are high voltage low leakage capacitors having a combined energy loss to leakage on the order of tens of wW. Low leakage rate capacitors 56 and 58 are chosen so that they can be fully or partially charged and then retain a substantial part of that charge over a relatively extended period of time.
In this embodiment batteries 11 and 13 are used to charge capacitors 56 and 58 as discussed above. Preferably batteries 11 and 13 are medium rate batteries or a two cell combination of a low rate and high rate battery. However, as stated above, it is fully contemplated that any combination or any type of battery including a single battery could be used without departing from the spirit of the invention. The medium rate battery is smaller in size compared to a high rate battery and thus volume within the implantable device can be significantly reduced. Nevertheless, it is fully contemplated that a high rate battery could be utilized within the implantable device to charge capacitors 56 and 58. However, with a medium rate battery, capacitors 56 and 58 can be charged over a relatively long time, such as between 20 seconds to several minutes.
As stated above this is better for batteries 11 and 13 and will increase theix lifetime, which thus increases the implantable device's lifetime. Further, since it is also desirable to minimize the volume occupied by the implantable devices as well as their mass to further limit patient discomfort, a smaller medium rate battery is preferred.
With reference to Figures 4 and 5, a flow diagram of an embodiment for capacitor optimization and a table of an embodiment for capacitor optimization is shown.
In the present embodiment, batteries 11 and 13 first charge capacitors 56 and 58 to an initial level, which is shown as state 400. Preferably capacitors 56 and 58 are fully charged as represented by region 500 of figure 5, however, it is contemplated that capacitors 56 arid 58 could be partially charged, which would require a shorter charging time upon detection of an arrhythmia, and thus a shorter time until a therapeutic shock could be delivered. Microprocessor 42 continuously receives the VCAP signal giving the voltage levels of capacitors 56 and 58 from control circuitry 40 via data bus 46.
Processor 42 monitors the voltage level of capacitors 56 and 58 and determines if the capacitor is fully charged, as is shown in state 402. If capacitors 56 and 58 axe not IO fully charged, processor 42 maintains charging of capacitors 56 and 58, thus returning to state 400. However, if capacitors 56 and 58 are fully charged, then processor 42 creates an open circuit between batteries 11 and 13 and capacitors 56 and 58, as shown in state 404. It is contemplated that processor 42 could create this open circuit by opening a relay switch, turning on or off a transistor, or utilizing any other switching 15 methods known in the art.
Processor 42 then determines from the VCAP signal whether capacitors 56 and 58 have fallen below a predetermined charge, as shown in state 406. Preferably this predetermined level is chosen during implantation of the implantable medical device and is chosen to be a level, which can provide an adequate shock to correct an 20 arrhythmia. Over a period of hundreds of minutes, low leakage capacitors 56 and 58 will eventually loose enough charge through current leakage that their charge will fall to a predetermined level represented by xegion 502 in Figure 5. When the charge level in capacitors 56 and 58 falls below this predeterniined level, microprocessor instructs confirol circuitry 40 to begin charging capacitors 56 and 58 as represented by 25 region 504. Thus processor 42 returns to state 400. If capacitors 56 and 58 have not fallen below the predetermined level, processor 42 determines whether an arrhythmia has been detected, shown as state 407 in Figure 4. If no arrhythmia is detected than processor 42 returns to state 404 to assure that batteries 11 and 13 are isolated from capacitors 56 and 58. If an arrhythmia is detected, processor 42 delivexs a therapeutic 30 shock at the quickest possible moment, as shown in state 408. It is well known that the shock cannot be delivered during certain times, therefore, the shock is delivered at the quickest possible moment. As stated above, this quickly delivered therapy substantially increases the efficacy of the therapy.
After the therapeutic shock is delivered, processor 42 returns to state 400 where batteries 11 and 13 are reconnected with capacitors 56 and 58 and begin charging them. Thus the remaining charge left after the therapy is not lost, since capacitors 56 and 58 quickly begin recharging after the therapy. Once capacitors 56 and 58 are fully charged again (state 402), processor 42 then instructs control 40 to stop charging capacitors 56 and 58 (state 404). This process then repeats continuously until an arrhythmia is detected (state 407) in which case, as described above, capacitors 56 and 58 are discharged to provide a properly timed shock to the heart (state 408).
After a shock event, the present embodiment is preferably implemented so that the total time to second shock is approximately 30 seconds. As is known, sometimes the first shock event is unsuccessful in stopping an arrhythmia; therefore, a second shock event is sometimes needed. The present invention is still able to supply a second shock in plenty of time even though a medium rate battery is being implemented. In the alternative, a high voltage binary battery could be implemented where if a second shock event was necessary, the binary battery would provide a high voltage charge to capacitors 56 and 58 within 5 to 20 seconds.
With reference to Figure 6, a flow diagram of an embodiment for capacitor optimization is shown. In this embodiment, batteries 11 and 13 first charge capacitors 56 and 58 to an initial level, which is shown as state 600. Processor 42 monitors the voltage level of capacitors 56 and 58 and determines if the capacitor is fully charged, as is shown in state 602. If capacitors 56 and 58 are not fully charged, processor 42 maintains the charging of capacitors 56 and 58, thus returning to state 600.
However, if capacitors 56 and 58 are fully charged, then processor 42 creates an open circuit between batteries 11 and 13 and capacitors 56 and 58, as shown in state 604.
It is contemplated that processor 42 could create this open circuit by opening a relay switch, turning off or on a transistor, or utilizing any other switching methods known in the art.
Processor 42 then determines whether a predetermined amount of time has expired since capacitors 56 and 58 were fully charged as represented by state 606.
Preferably this predeterniined time period represents the time it takes before the leakage current of capacitors 56 and 58 have drained the charge on capacitors 56 and 58 to a level just above one which could provide a shock to correct an arrhythmia event. Over a period of hundreds of minutes, low leakage capacitors 56 and 58 will eventually loose enough charge through current leakage that their charge will fall below an effective charge.
When this predetermined time period has passed, microprocessor 42 instructs control circuitry 40 to begin charging capacitors 56 and 58 as represented by region 604. Thus processor 42 returns to state 600. If the predetermined time period has not passed, processor 42 determines whether an arrhythmia has been detected, shown as state 607.
If no arrhythmia is detected than processor 42 returns to state 606 to determine whether the predetermined time limit has passed. If an arrhythmia is detected, processor 42 delivers a therapeutic shock at the quickest possible moment, as shown in I S state 608. After the therapeutic shock is delivered, processor 42 returns to state 600 where batteries 11 and 13 are reconnected with capacitors 56 and 58 and begin charging them.
In another embodiment, batteries 11 and 13 supply a continual medium rate charge to capacitors 56 and 58 to maintain them at a full or partial charge. In this embodiment, once capacitors 56 and 58 are preferably at maximum charge batteries 11 and 13 only have to supply capacitors 11 and 13 with enough charge to replace the charge lost due to the leakage current in order to keep capacitors 56 and 58 at a substantially full charge. Since the leakage current is so low for capacitors 56 and 58, the amount of charge required from batteries 11 and 13 is low. Thus, the continual charging does not deplete batteries 11 and 13. In comparison the leakage current of capacitors 56 and 58 is lower than the current required by processor 42. In this embodiment, capacitors 56 and 58 preferably don't fall below a full charge.
Similar to above, when an arrhythmia is detected, a shock can be delivered at the quickest possible moment thus increasing the efficacy of the shock and the more likely normal cardiac rhythm is successfully restored.
In another embodiment batteries 11 and 13 charge a low leakage capacitor, which in turn charges a high voltage capacitor. In this embodiment, capacitors 56 and 58 could be any type of capacitors and would not have to be low leakage capacitors.
Batteries 11 and 13 would continuously charge the Iow leakage capacitor and in the event of an arrhythmia, the low leakage capacitor would discharge through transformer 68, thus almost instantly charging capacitors 56 and 58, which would discharge immediately upon reaching full charge. It is noted that the Iow leakage capacitor retains any charge not delivered to capacitors 56 and 58 so that no charge is wasted. It is also contemplated that any combination of the embodiments listed above could be utilized without departing from the spirit of the invention.
In another embodiment a binary, chemical or thermal battery is utilized to power a "lifeboat" type of defibrillator. This device would be essentially inactive except for a monitoring circuit, such as in a pacemaker until an arrhythmia was detected.
Upon detection, the binary (or thermal, chemical) battery would be activated and provide a high voltage shock at the quickest possible moment.
It will be appreciated that the present invention can take many forms and embodiments. The true essence and spirit of this invention are defined in the appended claims, and it is not intended that the embodiment of the invention presented herein should limit the scope thereof.
Therefore, it would be desirable to capture these remaining charges and thus extend the Iife of the battery.
For the foregoing reasons, there is a need for an ICD, which allows for a relatively long charging time and yet retains clinical efficacy to prolong battery life and provide for a smaller battery. There is also a need for an ICD providing a high voltage capacitor with very low leakage current so that the capacitor could be held at full charge thus reducing the adverse effects of rapid charging. There is also a need for an ICD that when an arrhythmia is detected the ICD can deliver therapy at the quickest possible moment without having to wait for a capacitor to charge thus increasing the efficacy of the delivered therapy.
A medical device for electrical termination of an arrhythmic condition of a patient's heart in embodiments of the invention may include one or more of the following features: (a) at Ieast one battery; (b) means for detection of an arrhythmic condition of a patient's heart; (c) at least one high voltage capacitor; (d) converter means for providing charging current from said at least one battery to said at least one capacitor; (e) means for maintenance of a charge on said at least one capacitor between arrhythmia therapies; (f) controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and (g) discharge circuit means for delivering voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
A method for electrical termination of an arrhythmic condition of a patient's heart in embodiments of the invention may include one or more of the following features:
(a) charging at Ieast one high voltage capacitor with current from at least one battery, (b) detecting an arrhythmic condition of a patient's heart, (c) maintaining the charge on said at least one capacitor between arrhythmia therapies, (d) providing a controller means responsive to detection of an arrhythmic condition of said patient's heart, (e) generating a discharge control signal upon detection of an arrhythmic condition of said patient's heart; and (f) delivering a voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
FIG. 1 is a drawing illustrating the general physical components of a pacemaker/cardioverter/defibrillator and lead system of the type in which the present invention may be advantageously practiced;
FIG. 2 is a functional block diagram illustrating the general interconnection of voltage conversion circuitry of the present invention with the primary functional components of an implantable pacemaker/cardioverter/defibrillator;
FIG. 3 is a schematic block diagram of the general components of a pacemaker/cardioverter/defibrillator employing a high voltage charging circuit;
FIG. 4 is a flow diagram of an embodiment for capacitor optimization of the present invention;
FIG. 5 is a table representing a capacitor optimization embodiment of the present invention;
FIG. 6 is a flow diagram of an embodiment for capacitor optimization of the present invention;
The following detailed description is to be read with reference to the figures, in which like elements in different figures have like reference numerals. The figures, which are not necessarily to scale, depict selected embodiments and are not intended to limit the scope of the invention. Skilled artisans will recognize that the examples provided herein have many useful alternatives that fall within the scope of the invention.
The pxesent invention is not limited to implantable cardioverter defibrillators and may be employed in many various types of electronic and mechanical devices for treating patient medical conditions such as external cardioverter defibrillators, pacemakers, and neurostimulators. It is to be further understood; moreover, the present invention is not limited to medium current rate batteries and may be utilized for low and high current rate batteries. For purposes of illustration only, however, the present invention is below described in the context of medium current rate batteries and implantable cardioverter defibrillators.
The present invention is described generally in a system providing biphasic cardioversion pulses or shocks in a cardioversion system. However, it is fully contemplated that the present invention could be utilized in any type of pulse or shock delivery methodology utilizing any type of pulse of shock waveform without departing from the spirit of the invention. In the description of the preferred embodiment that follows, an implantable pacemaker/cardioverter/defibrillator in which the present invention is preferably implemented is capable of providing monophasic, biphasic, or any other caxdioversion pulse or shock waveform. However, a variety of implantable leads and electrode systems may be employed, with more than one cardioversion electrode connected electrically in common to widen the cardioversion energy distribution across the heart. Such electrodes may include indwelling right ventricular, superior vena cava, and coronary sinus electrodes, active pulse generator case electrodes and/or epicardial and subcutaneous patch electrodes in various combinations of two or more. With a three electrode system, two of the electrodes are connected in common, and the energy distribution between the two common and the third electrode may lead to reduced energy sufficient to reliably cardiovert a heart in fibrillation or high rate malignant ventricular tachycardia.
FIG. 1 illustrates such a general implementation of an implantable pacemaker/cardioverter/defibrillator 10 and one possible selection of cardioversion electrodes on associated electrical leads 14, 16 and 18, and their relationship to a human heart 12. The leads 14, 16, and 18 are coupled to the pacemaker/cardioverter/defibrillator 10 by means of a mufti-port connector block 20, which contains separate connector ports fox each of the three leads illustrated. Each of the leads 14, 16, 18 comprise a large surface area cardioversion electrode, and lead 18 _7_ also comprises a pair of pace/sense electrodes (making it a tripolar lead) all as described below.
Unipolar lead 14 is coupled to a subcutaneous cardioversion electrode 30, which is intended to be mounted subcutaneously in the region of the left chest.
Unipolar lead 16 is a coronary sinus (CS) lead employing an elongated coil, cardiovexsion electrode that is located in the coronary sinus of the heart. When positioned in the CS, the CS
electrode extends around the heart from a point within the opening or ostium of the CS
to a point in the vicinity of the left atrial appendage, as shown in broken line format at 32.
Tripolar lead 18 is provided with an elongated electrode coil 28 which is located in the right ventricle of the heart and functions as a third cardioversion electrode. Lead 18 also includes a first pace/sense electrode 34 and a second, closely spaced, pace/sense electrode 38. Electrode 34 takes the form of a distal helical coil, which is screwed into the myocardial tissue of the right ventricle. The second pace/sense electrode 38 is closely spaced to the electrode 34 for bipolar pacing and near fteld electrogram or R-wave sensing in the apex of the right ventricle. A more detailed description of the leads illustrated can be found in U.S. Pat. No. 5,163,427, herein incorporated by reference in its entirety.
Through testing at implantation of cardioversion efficacy across one of the three electrodes with the other two electrodes in common or with each of the other electrodes alone, a selection may be made of the most efficacious electrode selection.
If only two electrodes are needed, then the third lead and electrode may be eliminated.
Typically, it is expected that all thxee of the electrodes will be employed, with two connected electrically in common internally within the pulse generator 10 as described below.
FIG. 2 is a block diagram illustrating the general interconnections of a voltage output circuit 40, a voltage charging circuit 64 and capacitor bank 56, 58 according to one embodiment of the present invention with a prior art implantable pacemakex/cardioverter/defibrillator. As illustrated, the device is controlled by means _g_ of a stored program in a microprocessor 42, which performs all necessary computational functions within the device. Microprocessor 42 is linked to control circuitry 44 by means of a bi-directional datalcontrol bus 46, and thereby controls operation of the output circuitry 40 and the high voltage charging circuitry 64. On reprogramming of the device or on the occurrence of signals indicative of delivery of cardiac pacing pulses or of the occurrence of cardiac contractions, pace/sense circuitry 78 will awaken microprocessor 42 to perform any necessary mathematical calculations, to perform tachycardia and fibrillation detection procedures and to update the time intervals controlled by the timers in pace/sense circuitry 78.
The control circuitry 44 provides three signals of primary importance to the output circuitry 40 of the present invention. These include the first and second control signals discussed above, labeled here as ENAB, line 48, and ENBA, line 50, which govern the timing and duration of the two phases of the biphasic cardioversion pulse or shock.
Also of importance is the DUMP signal on line 52, which initiates discharge of the output capacitors, and the VCAP signal on line 54, which is indicative of the voltage stored on the output capacitors C1, C2, and is applied to the control circuitry 44.
As described above, a wide variety of cardioversion electrode bearing leads may be attached to two or all three cardioversion output terminals, labeled HVX, HVA, and HVB in FIG. 2, coupled to the connector block 20 bores. In the example illustrated in FIGS. 1 and 2, it will be assumed that the electrodes 28, 30 and 32 are coupled to the high voltage output circuitry 40 by means of connectors in the connector block illustrated as conductors 22, 24 and 26, respectively. As shown in FIG. 3, conductors 22 and 24 labeled HVX and HVA are electrically connected in common so that an output shock may be delivered even if all three leads 18, 14 and 16 and electrodes 28, 30 and 32, respectively, are connected to the pulse generator as shown in FIG.
1 and described above.
The high voltage output circuit 40 includes a capacitor bank, including capacitors 56 and 58, which is discussed in more detail below, and diodes 70 and 72, used for delivering defibrillation pulses to the electrodes. In FIG. 2, the capacitor bank is illustrated in conjunction with the high voltage charging circuitry 64, controlled by the control/timing circuitry 44 by means of CHDR line 66. As illustrated, capacitors 56 (C1) and 58 (C2) axe charged by means of a high frequency, high voltage transformer 68. Proper charging polarities are maintained by means of the diodes 70 and 72.
VCAP line 54 provides a signal indicative of the voltage on the capacitor bank, and allows for control of the high voltage charging circuitry and for termination of the charging function when the stored voltage equals the programmed charging level.
The delivery of the biphasic cardioversion shock is controlled by the partial discharge of the voltage on the output capacitor bank in a first direction during a first phase logic signal on ENAB, line 48, and by further discharge of the remaining voltage in a second direction during closely timed second signal on ENBA, line 50.
When ENAB is present, the first phase of the cardioversion pulse is delivered between the electrodes) 30 and/or 32 and electrode 28. During a logic signal on ENBA, line 50, the second phase is delivered between in the opposite direction between the same electrodes.
Pace/sense circuitry 78 includes an R-wave amplifier according to the prior art, or more advantageously as disclosed in U.S. Pat. No. 5,117,824 by Keimel et al, which is incorporated herein by reference in its entirety. However, the present invention is believed workable in the context of any known R-wave amplification system.
Pace/sense circuitry 78 also includes a pulse generator for generating cardiac pacing pulses, which may also correspond to any known cardiac pacemaker output circuitry and includes timing circuitry for def ning ventricular pacing intervals, refractory intervals and blanking intervals, under control of microprocessor 42 via control/data bus 80. Control signals triggering generation of cardiac pacing pulses by pace/sense circuitry 78 and signals indicative of the occurrence of R-waves, from pace/sense circuitry 78 are communicated to control circuitry 44 by means of a bi-directional data bus 81. Pace/sense circuitry 78 is coupled to helical electrode 34 and ring electrode 38 of tripolar lead 18 through connector elements of the connector block 20 and associated adapters, if necessary, illustrated schematically as conductors 36 and 37.
The present invention constitutes an apparatus and method for maintaining a full or partial charge on a capacitor within an implantable medical device between therapies.
The particular circuitry or components involved in the implementation of shock timing optimization axe shown in specific detail. However, it is fully contemplated that alternate circuitry or components could be utilized, such as described in U.S. Pat.
No. 6,438,420 (Thompson) herein incorporated by reference, without departing from the spirit of the invention. A number of additional expressions for input and output signals or terminals than those described above axe used throughout, including:
CHGDR--Charge drive signal for driving the on/off switch in the primary winding of the flyback transformer at a duty cycle established by the relative on and off times.
VSS--VSS is the circuit ground, which may also appear labeled QVSS and may be connected to BATTN.
BATT--Battery positive power supply, which may also appear as B+ or as BP.
BATTN--Battery negative power supply.
PPLUS--Plus terminal for the pace/sense function.
PMINUS--Negative terminal for the pace/sense function.
ENBA--Enable signal commanding capacitor discharge from HVB to HVA (and HVX) and setting the duration of one phase of the biphasic pulse.
ENAB--Enable signal commanding capacitor discharge from HVA (and HVX) to HVB and setting the duration of the other phase of the biphasic pulse.
CSP--Charge store positive terminal.
C1P--Capacitor 1 positive terminal connection.
C1N--Capacitor 1 negative terminal connection.
C2P--Capacitor 2 positive terminal connection.
C2N--Capacitor 2 negative terminal connection.
-lI-CSN--Charge store negative terminal.
VDD--Internally generated programmable regulated power supply.
DUMP--DUMP signal initiates the internal self discharge of the capacitoxs C1, to a load impedance.
OPTIN--Input terminal to the drive circuit optionally connected to an opto-coupler.
VIN--Input terminal to the drive circuit optionally connected to an input signal source.
VOUT--Output terminal of the drive circuit for supplying VDD voltage.
CSEN--Enable signal input terminal of the drive circuit optionally coupled to receive an opto-coupler command signal.
CSOUT--Output terminal of the drive circuit optionally coupled to drive an opto-coupler.
Other acronyms may appear in the description of the following drawings, which will be explained as necessary to understand the manner in which the present invention may be practiced in its preferred embodiment.
Turning now to FIG. 3, the circuit components of the pacemakerlcardioverter/defibrillator of the present invention are depicted and they include the batteries 11 and 13, the PC board 102, the high voltage output capacitors C1, C2 (56, 58 in FIG. 2), the high power hybrid board 104, the low power hybrid boaxd 106, the crystal 15, the antenna 17, and the reed switch 19. The batteries 11 and 13 are coupled to the BATT and BATTN inputs of the PC board 102. Although two batteries are shown, it is fully contemplated that any type or combination of batteries could be utilized, such as a single cell battery, a dual cell battery, or a mixture of high current and low current cells, without departing from the spirit of the invention. The crystal 15 is coupled to the Xl and X1N inputs of the low power hybrid 106.
The antenna 17 is coupled between the ANT and ANTGND inputs of low power hybrid 106 and the reed switch 19 is coupled between the RDSW and RSGND inputs of low power hybrid 106. The PPLUS and PMINLTS terminals are coupled to respectively labeled pins of the low power hybrid 106, which contains the pace/sense circuitry 78 of FIG. 2.
The low power hybrid 106 includes the basic timing and control circuitry of the system, including the programming and telemetry functions, the electrogram sensing and pacing functions, the microprocessor and RAM/ROM memories, all implemented in both digital and analog circuits corresponding to blocks 42, 44 and 78 in FIG. 2.
The low power hybrid 106 develops the CHGDR signal as well as the DUMP, ENBA
and ENAB signals relevant to the operation of the high voltage output circuit of the present invention.
The PC board 102 corresponds to the high voltage-charging block 64 in FIG. 2, and also includes the step up transformer 110 and diodes 121, 123. The relatively large output capacitors C1, C2 are electrically connected to the PC boaxd 102 through the input terminals C1N and C1P and C2N and C2P, respectively. The PC board presents the charge storage positive and negative signals CSP and CSN, respectively, to the high power hybrid 104. PC board 102 also includes an on-off control switch, responsive to the CHGDR signal from the low power hybrid 106, for supplying stepped up, rectified current to the output capacitors C1, C2, across which the voltage signals CSP, CSN are developed.
The high power hybrid 104 corresponds to the high voltage output block 40 illustrated in FIG. 2 and includes switching circuitry for delivery of voltage stored in capacitors C1 and C2 as monophasic, biphasic, or any other output pulse waveform.
Delivery of the output pulses is controlled by the low power hybxid 106 via ENAB and ENBA lines 48 and 50, respectively. Similarly, the HVA line 24, which is coupled in common to the HVX line 22, and the HVB line 26 are coupled to the HVA and HVB
output pins of high power hybrid 104. The high voltage discharges forming the cardioversion shocks are generated from the high power hybrid 104 and conducted to the HVA and HVB output terminals and the cardioversion electrode system employed as described above.
With reference to Figures 2 and 3 again, one embodiment of the present invention is described. In one embodiment, capacitors 56 and 58 are high voltage capacitors with an extremely low leakage current. An exemplary low leakage capacitor is described in U.S. Pat. No. 5,808,856 (Bischoff, et. al.), U.S. Pat. No.
6,426,863 (Munshi) and U.S. Pub. No. 2002/0052078 (Zheng et. al.). While it is preferable that a low leakage capacitor be utilized for the present embodiment, it is contemplated that any high voltage capacitor could be utilized without departing from the spirit of the invention. Further, it is fully contemplated that the present invention could utilize one or more individual capacitors as well as multiple capacitors utilizing a wide range of capacitor voltages. Nevertheless, preferably capacitors 56 and 58 are high voltage low leakage capacitors having a combined energy loss to leakage on the order of tens of wW. Low leakage rate capacitors 56 and 58 are chosen so that they can be fully or partially charged and then retain a substantial part of that charge over a relatively extended period of time.
In this embodiment batteries 11 and 13 are used to charge capacitors 56 and 58 as discussed above. Preferably batteries 11 and 13 are medium rate batteries or a two cell combination of a low rate and high rate battery. However, as stated above, it is fully contemplated that any combination or any type of battery including a single battery could be used without departing from the spirit of the invention. The medium rate battery is smaller in size compared to a high rate battery and thus volume within the implantable device can be significantly reduced. Nevertheless, it is fully contemplated that a high rate battery could be utilized within the implantable device to charge capacitors 56 and 58. However, with a medium rate battery, capacitors 56 and 58 can be charged over a relatively long time, such as between 20 seconds to several minutes.
As stated above this is better for batteries 11 and 13 and will increase theix lifetime, which thus increases the implantable device's lifetime. Further, since it is also desirable to minimize the volume occupied by the implantable devices as well as their mass to further limit patient discomfort, a smaller medium rate battery is preferred.
With reference to Figures 4 and 5, a flow diagram of an embodiment for capacitor optimization and a table of an embodiment for capacitor optimization is shown.
In the present embodiment, batteries 11 and 13 first charge capacitors 56 and 58 to an initial level, which is shown as state 400. Preferably capacitors 56 and 58 are fully charged as represented by region 500 of figure 5, however, it is contemplated that capacitors 56 arid 58 could be partially charged, which would require a shorter charging time upon detection of an arrhythmia, and thus a shorter time until a therapeutic shock could be delivered. Microprocessor 42 continuously receives the VCAP signal giving the voltage levels of capacitors 56 and 58 from control circuitry 40 via data bus 46.
Processor 42 monitors the voltage level of capacitors 56 and 58 and determines if the capacitor is fully charged, as is shown in state 402. If capacitors 56 and 58 axe not IO fully charged, processor 42 maintains charging of capacitors 56 and 58, thus returning to state 400. However, if capacitors 56 and 58 are fully charged, then processor 42 creates an open circuit between batteries 11 and 13 and capacitors 56 and 58, as shown in state 404. It is contemplated that processor 42 could create this open circuit by opening a relay switch, turning on or off a transistor, or utilizing any other switching 15 methods known in the art.
Processor 42 then determines from the VCAP signal whether capacitors 56 and 58 have fallen below a predetermined charge, as shown in state 406. Preferably this predetermined level is chosen during implantation of the implantable medical device and is chosen to be a level, which can provide an adequate shock to correct an 20 arrhythmia. Over a period of hundreds of minutes, low leakage capacitors 56 and 58 will eventually loose enough charge through current leakage that their charge will fall to a predetermined level represented by xegion 502 in Figure 5. When the charge level in capacitors 56 and 58 falls below this predeterniined level, microprocessor instructs confirol circuitry 40 to begin charging capacitors 56 and 58 as represented by 25 region 504. Thus processor 42 returns to state 400. If capacitors 56 and 58 have not fallen below the predetermined level, processor 42 determines whether an arrhythmia has been detected, shown as state 407 in Figure 4. If no arrhythmia is detected than processor 42 returns to state 404 to assure that batteries 11 and 13 are isolated from capacitors 56 and 58. If an arrhythmia is detected, processor 42 delivexs a therapeutic 30 shock at the quickest possible moment, as shown in state 408. It is well known that the shock cannot be delivered during certain times, therefore, the shock is delivered at the quickest possible moment. As stated above, this quickly delivered therapy substantially increases the efficacy of the therapy.
After the therapeutic shock is delivered, processor 42 returns to state 400 where batteries 11 and 13 are reconnected with capacitors 56 and 58 and begin charging them. Thus the remaining charge left after the therapy is not lost, since capacitors 56 and 58 quickly begin recharging after the therapy. Once capacitors 56 and 58 are fully charged again (state 402), processor 42 then instructs control 40 to stop charging capacitors 56 and 58 (state 404). This process then repeats continuously until an arrhythmia is detected (state 407) in which case, as described above, capacitors 56 and 58 are discharged to provide a properly timed shock to the heart (state 408).
After a shock event, the present embodiment is preferably implemented so that the total time to second shock is approximately 30 seconds. As is known, sometimes the first shock event is unsuccessful in stopping an arrhythmia; therefore, a second shock event is sometimes needed. The present invention is still able to supply a second shock in plenty of time even though a medium rate battery is being implemented. In the alternative, a high voltage binary battery could be implemented where if a second shock event was necessary, the binary battery would provide a high voltage charge to capacitors 56 and 58 within 5 to 20 seconds.
With reference to Figure 6, a flow diagram of an embodiment for capacitor optimization is shown. In this embodiment, batteries 11 and 13 first charge capacitors 56 and 58 to an initial level, which is shown as state 600. Processor 42 monitors the voltage level of capacitors 56 and 58 and determines if the capacitor is fully charged, as is shown in state 602. If capacitors 56 and 58 are not fully charged, processor 42 maintains the charging of capacitors 56 and 58, thus returning to state 600.
However, if capacitors 56 and 58 are fully charged, then processor 42 creates an open circuit between batteries 11 and 13 and capacitors 56 and 58, as shown in state 604.
It is contemplated that processor 42 could create this open circuit by opening a relay switch, turning off or on a transistor, or utilizing any other switching methods known in the art.
Processor 42 then determines whether a predetermined amount of time has expired since capacitors 56 and 58 were fully charged as represented by state 606.
Preferably this predeterniined time period represents the time it takes before the leakage current of capacitors 56 and 58 have drained the charge on capacitors 56 and 58 to a level just above one which could provide a shock to correct an arrhythmia event. Over a period of hundreds of minutes, low leakage capacitors 56 and 58 will eventually loose enough charge through current leakage that their charge will fall below an effective charge.
When this predetermined time period has passed, microprocessor 42 instructs control circuitry 40 to begin charging capacitors 56 and 58 as represented by region 604. Thus processor 42 returns to state 600. If the predetermined time period has not passed, processor 42 determines whether an arrhythmia has been detected, shown as state 607.
If no arrhythmia is detected than processor 42 returns to state 606 to determine whether the predetermined time limit has passed. If an arrhythmia is detected, processor 42 delivers a therapeutic shock at the quickest possible moment, as shown in I S state 608. After the therapeutic shock is delivered, processor 42 returns to state 600 where batteries 11 and 13 are reconnected with capacitors 56 and 58 and begin charging them.
In another embodiment, batteries 11 and 13 supply a continual medium rate charge to capacitors 56 and 58 to maintain them at a full or partial charge. In this embodiment, once capacitors 56 and 58 are preferably at maximum charge batteries 11 and 13 only have to supply capacitors 11 and 13 with enough charge to replace the charge lost due to the leakage current in order to keep capacitors 56 and 58 at a substantially full charge. Since the leakage current is so low for capacitors 56 and 58, the amount of charge required from batteries 11 and 13 is low. Thus, the continual charging does not deplete batteries 11 and 13. In comparison the leakage current of capacitors 56 and 58 is lower than the current required by processor 42. In this embodiment, capacitors 56 and 58 preferably don't fall below a full charge.
Similar to above, when an arrhythmia is detected, a shock can be delivered at the quickest possible moment thus increasing the efficacy of the shock and the more likely normal cardiac rhythm is successfully restored.
In another embodiment batteries 11 and 13 charge a low leakage capacitor, which in turn charges a high voltage capacitor. In this embodiment, capacitors 56 and 58 could be any type of capacitors and would not have to be low leakage capacitors.
Batteries 11 and 13 would continuously charge the Iow leakage capacitor and in the event of an arrhythmia, the low leakage capacitor would discharge through transformer 68, thus almost instantly charging capacitors 56 and 58, which would discharge immediately upon reaching full charge. It is noted that the Iow leakage capacitor retains any charge not delivered to capacitors 56 and 58 so that no charge is wasted. It is also contemplated that any combination of the embodiments listed above could be utilized without departing from the spirit of the invention.
In another embodiment a binary, chemical or thermal battery is utilized to power a "lifeboat" type of defibrillator. This device would be essentially inactive except for a monitoring circuit, such as in a pacemaker until an arrhythmia was detected.
Upon detection, the binary (or thermal, chemical) battery would be activated and provide a high voltage shock at the quickest possible moment.
It will be appreciated that the present invention can take many forms and embodiments. The true essence and spirit of this invention are defined in the appended claims, and it is not intended that the embodiment of the invention presented herein should limit the scope thereof.
Claims (37)
1. A medical device comprising:
at least one battery;
means for detection of an arrhythmic condition;
at least one high voltage capacitor;
converter means for providing charging current from said at least one battery to said at least one capacitor;
means for maintenance of a charge on said capacitor between arrhythmia therapies;
controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and discharge circuit means for delivering voltage stored on said at least one capacitor to said patient's heart in response to said discharge control signal.
at least one battery;
means for detection of an arrhythmic condition;
at least one high voltage capacitor;
converter means for providing charging current from said at least one battery to said at least one capacitor;
means for maintenance of a charge on said capacitor between arrhythmia therapies;
controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and discharge circuit means for delivering voltage stored on said at least one capacitor to said patient's heart in response to said discharge control signal.
2. The medical device of claim 1, wherein the at least one high voltage capacitor is a low leakage high voltage capacitor.
3. The medical device of claim 1, wherein the at least one battery provides for a capacitor charging time greater than about 10 seconds.
4. The medical device of claim 3, wherein the charging time provides for a reduction in battery size.
5. The medical device of claim 4, wherein the reduction in battery size provides for a reduction in medical device size.
6. The medical device of claim 1, wherein the voltage stored on said capacitor is delivered to said patient's heart at the quickest appropriate moment upon the detection of an arrhythmia condition of a patient's heart.
7. The medical device of claim 1, wherein the means for maintenance is comprised of an isolation means for disconnecting the converting means from said at least one capacitor after a predetermined charge level is reached on said at least one capacitor.
8. The medical device of claim 7, wherein the converting means is isolated from the at least one capacitor for a predetermined amount of time before the converting means is reconnected to the at least one capacitor to restore any charge lost due to leakage.
9. A medical device for electrical termination of an arrhythmic condition of a patient's heart of the type comprising:
at least one battery;
means for detection of an arrhythmic condition of a patient's heart;
at least one high voltage capacitor having low current leakage, said capacitor maintaining a charge between arrhythmia therapies;
converter means for providing charging current from said battery to said capacitor;
controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and discharge circuit means for delivering voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
at least one battery;
means for detection of an arrhythmic condition of a patient's heart;
at least one high voltage capacitor having low current leakage, said capacitor maintaining a charge between arrhythmia therapies;
converter means for providing charging current from said battery to said capacitor;
controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and discharge circuit means for delivering voltage stored on said capacitor to said patient's heart in response to said discharge control signal.
10. The medical device of claim 9, wherein the at least one battery provides for a capacitor charging time greater than about 20 seconds.
11. The medical device of claim 9, wherein the at least one battery provides for a capacitor charging time greater than about 10 seconds.
12. The medical device of claim 11, wherein the charging time provides for a reduction in battery size.
13. The medical device of claim 12, wherein the reduction in battery size provides for a reduction in medical device size.
14. The medical device of claim 9, wherein the voltage stored on said at least one capacitor is delivered to said patient's heart at the quickest appropriate moment upon the detection of an arrhythmia condition of a patient's heart.
15. The medical device of claim 14, further comprising an isolation means for disconnecting the converting means from said capacitor after a predetermined charge level is reached on said capacitor.
16. The medical device of claim 15, wherein the converting means is isolated from the capacitor for a predetermined amount of time before the converting means is reconnected to the capacitor to restore any charge lost due to leakage.
17. The medical device of claim 9, wherein the capacitor is maintained at a full charge between arrhythmia therapies.
18. The medical device of claim 9, wherein the capacitor is maintained at a partial charge between arrhythmia therapies.
19. A medical device for electrical termination of an arrhythmic condition of a patient's heart of the type comprising:
at least one battery;
means for detection of an arrhythmic condition of a patient's heart;
at least one high voltage charge capacitor having low current leakage, said at least one high voltage charge capacitor maintains a charge between arrhythmia therapies;
at least one high voltage delivery capacitor;
converter means for providing charging current from said battery to said charge capacitor;
controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and discharge circuit means for delivering voltage stored on said charge capacitor quickly to delivery capacitor and then delivering voltage stored on said delivery capacitor to said patient's heart in response to said discharge control signal.
at least one battery;
means for detection of an arrhythmic condition of a patient's heart;
at least one high voltage charge capacitor having low current leakage, said at least one high voltage charge capacitor maintains a charge between arrhythmia therapies;
at least one high voltage delivery capacitor;
converter means for providing charging current from said battery to said charge capacitor;
controller means responsive to detection of an arrhythmic condition of said patient's heart and for providing a discharge control signal; and discharge circuit means for delivering voltage stored on said charge capacitor quickly to delivery capacitor and then delivering voltage stored on said delivery capacitor to said patient's heart in response to said discharge control signal.
20. The medical device of claim 19, wherein the battery provides for a capacitor charging time greater than about 20 seconds.
21.The medical device of claim 20, wherein the battery provides for a capacitor charging time greater than about 10 seconds.
22. The medical device of claim 21, wherein the charging time provides for a reduction in battery size.
23. The medical device of claim 19, wherein the reduction in battery size provides for a reduction in medical device size.
24.The medical device of claim 19, wherein the voltage stored on said delivery capacitor is delivered at a proper time to said patient's heart immediately upon the detection of an arrhythmia condition of a patient's heart.
25. The medical device of claim 24, further comprising an isolation means for disconnecting the converting means from said charging capacitor after a predetermined charge level is reached on said charging capacitor.
26. The medical device of claim 25, wherein the converting means is isolated from the charging capacitor for a predetermined amount of time before the converting means is reconnected to the charging capacitor to restore any charge lost due to leakage.
27. The medical device of claim 19, wherein the charging capacitor is maintained at a full charge between arrhythmia therapies.
28. The medical device of claim 19, wherein the charging capacitor is maintained at a
29. A method for electrical termination of an arrhythmic condition of a patient's heart of the type comprising the steps of:
charging at least one high voltage capacitor with current from at least one battery;
detecting an arrhythmic condition of a patient's heart;
maintaining the charge on said at least one capacitor between arrhythmia therapies;
providing a controller means responsive to detection of an arrhythmic condition of said patient's heart;
generating a discharge control signal upon detection of an arrhythmic condition of said patient's heart; and delivering a voltage stored on said at least one capacitor to said patient's heart in response to said discharge control signal.
charging at least one high voltage capacitor with current from at least one battery;
detecting an arrhythmic condition of a patient's heart;
maintaining the charge on said at least one capacitor between arrhythmia therapies;
providing a controller means responsive to detection of an arrhythmic condition of said patient's heart;
generating a discharge control signal upon detection of an arrhythmic condition of said patient's heart; and delivering a voltage stored on said at least one capacitor to said patient's heart in response to said discharge control signal.
30.The method of claim 29, wherein the at least one high voltage capacitor is a low leakage high voltage capacitor.
31. The method of claim 29, wherein the at least one battery provides for a capacitor charging time greater than about 10 seconds.
32. The method of claim 29, wherein the at least one battery provides for a capacitor charging time greater than about 10 seconds.
33. The method of claim 32, wherein the charging time provides for a reduction in battery size.
34. The method of claim 33, wherein the reduction in battery size provides for a reduction in medical device size.
35. The method of claim 34, wherein the voltage stored on said at least one capacitor is delivered to said patient's heart at the quickest appropriate moment upon the detection of an arrhythmia condition of a patient's heart.
36. The method of claim 35, further comprising the step of disconnecting the at least one battery from said capacitor after a predetermined charge level is reached on said at least one capacitor.
37. The method of claim 36, further comprising the step of isolating the at least one capacitor for a predetermined amount of time before reconnecting the at least one capacitor to the at least one battery to restore any charge lost due to leakage.
Applications Claiming Priority (3)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| US10/260,488 US20040064154A1 (en) | 2002-09-30 | 2002-09-30 | Apparatus and method for optimizing capacitor charge in a medical device |
| US10/260,488 | 2002-09-30 | ||
| PCT/US2003/030568 WO2004030749A1 (en) | 2002-09-30 | 2003-09-26 | Apparatus and method for optimizing capacitor charge in a medical device |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA2499842A1 true CA2499842A1 (en) | 2004-04-15 |
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| CA002499842A Abandoned CA2499842A1 (en) | 2002-09-30 | 2003-09-26 | Apparatus and method for optimizing capacitor charge in a medical device |
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| US (2) | US20040064154A1 (en) |
| EP (1) | EP1558329A1 (en) |
| CA (1) | CA2499842A1 (en) |
| WO (1) | WO2004030749A1 (en) |
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| WO2004030749A1 (en) | 2004-04-15 |
| US20060195148A1 (en) | 2006-08-31 |
| EP1558329A1 (en) | 2005-08-03 |
| US20040064154A1 (en) | 2004-04-01 |
| US8195291B2 (en) | 2012-06-05 |
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| Date | Code | Title | Description |
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| FZDE | Discontinued |